Volume 5 Supplement 1

Abstracts from the International Society for Therapeutic Ultrasound Conference 2016

Open Access

International Society for Therapeutic Ultrasound Conference 2016

Tel Aviv, Israel. 14-18 March, 2016
  • Brian Fowlkes1,
  • Pejman Ghanouni2,
  • Narendra Sanghvi3,
  • Constantin Coussios4,
  • Paul C. Lyon4,
  • Michael Gray4,
  • Christophoros Mannaris4,
  • Marie de Saint Victor4,
  • Eleanor Stride4,
  • Robin Cleveland4,
  • Robert Carlisle4,
  • Feng Wu7,
  • Mark Middleton6,
  • Fergus Gleeson5,
  • Jean-Franҫois Aubry8,
  • Kim Butts Pauly9,
  • Chrit Moonen10,
  • Jacob Vortman11,
  • Pejman Ghanouni12,
  • Shirley Sharabi13, 14,
  • Dianne Daniels13, 14,
  • David Last13,
  • David Guez13,
  • Yoav Levy15,
  • Alexander Volovick15,
  • Javier Grinfeld15,
  • Itay Rachmilevich15,
  • Talia Amar15,
  • Zion Zibly13,
  • Yael Mardor13, 14,
  • Sagi Harnof13, 14,
  • Michael Plaksin16,
  • Yoni Weissler16,
  • Shy Shoham16,
  • Eitan Kimmel16,
  • Omer Naor17, 18,
  • Nairouz Farah19,
  • Shy Shoham18,
  • Dong-Guk Paeng20, 22,
  • Zhiyuan Xu21,
  • John Snell20,
  • Anders H. Quigg20,
  • Matthew Eames20,
  • Changzhu Jin22,
  • Ashli C. Everstine23,
  • Jason P. Sheehan21,
  • Beatriz S. Lopes24,
  • Neal Kassell20,
  • Thomas Looi25, 28,
  • Vera Khokhlova26,
  • Charles Mougenot29,
  • Kullervo Hynynen27,
  • James Drake25,
  • Michael Slayton30,
  • Richard C. Amodei30,
  • Keegan Compton30,
  • Ashley McNelly31,
  • Daniel Latt31,
  • Michael Slayton32,
  • Richard C. Amodei32,
  • Keegan Compton32,
  • John Kearney33,
  • David Melodelima34, 35,
  • Aurelien Dupre35,
  • Yao Chen35,
  • David Perol35,
  • Jeremy Vincenot34,
  • Jean-Yves Chapelon34,
  • Michel Rivoire35, 34,
  • Wei Guo36,
  • Guoxin Ren36,
  • Guofeng Shen36,
  • Michael Neidrauer37,
  • Leonid Zubkov37,
  • Michael S. Weingarten38,
  • David J. Margolis39,
  • Peter A. Lewin37,
  • Nathan McDannold40,
  • Jonathan Sutton40,
  • Natalia Vykhodtseva40,
  • Margaret Livingstone41,
  • Thiele Kobus42, 43,
  • Yong-Zhi Zhang43,
  • Natalia Vykhodtseva43,
  • Nathan McDannold43,
  • Michael Schwartz44, 45,
  • Yuexi Huang45,
  • Nir Lipsman44,
  • Jennifer Jain45,
  • Martin Chapman45,
  • Tejas Sankar46,
  • Andres Lozano44,
  • Kullervo Hynynen45,
  • Michael Schwartz47, 48,
  • Robert Yeung47,
  • Yuexi Huang47,
  • Nir Lipsman48,
  • Jennifer Jain47,
  • Martin Chapman47,
  • Andres Lozano48,
  • Kullervo Hynynen47,
  • Christakis Damianou49,
  • Nikolaos Papadopoulos50,
  • Alexander Volovick51,
  • Javier Grinfeld51,
  • Yoav Levy51,
  • Omer Brokman51,
  • Eyal Zadicario51,
  • Ori Brenner52,
  • David Castel53,
  • Shih-Ying Wu54,
  • Julien Grondin54,
  • Wenlan Zheng54,
  • Marc Heidmann54,
  • Maria Eleni Karakatsani54,
  • Carlos J. Sierra Sánchez54,
  • Vincent Ferrera56,
  • Elisa E. Konofagou54, 55,
  • Christakis Damianou57,
  • Marinos Yiannakou57,
  • HongSeok Cho58,
  • Hwayoun Lee58,
  • Mun Han59,
  • Jong-Ryul Choi58,
  • Taekwan Lee58,
  • Sanghyun Ahn58,
  • Yongmin Chang59,
  • Juyoung Park58,
  • Nicholas Ellens60,
  • Ari Partanen60, 61,
  • Keyvan Farahani60, 62,
  • Raag Airan60,
  • Alexandre Carpentier65, 66,
  • Michael Canney63,
  • Alexandre Vignot63,
  • Cyril Lafon64,
  • Jean-Yves Chapelon64,
  • Jean-yves Delattre66, 67,
  • Ahmed Idbaih67,
  • Henrik Odéen68,
  • Bradley Bolster68,
  • Eun Kee Jeong69,
  • Dennis L. Parker69,
  • Pooja Gaur70,
  • Xue Feng71,
  • Samuel Fielden71,
  • Craig Meyer71,
  • Beat Werner72,
  • William Grissom70,
  • Michael Marx73,
  • Pejman Ghanouni73,
  • Kim Butts Pauly73,
  • Hans Weber74,
  • Valentina Taviani74,
  • Kim Butts Pauly74,
  • Pejman Ghanouni74,
  • Brian Hargreaves74,
  • Jun Tanaka75,
  • Kentaro Kikuchi75,
  • Ayumu Ishijima75,
  • Takashi Azuma75,
  • Kosuke Minamihata75,
  • Satoshi Yamaguchi75,
  • Teruyuki Nagamune75,
  • Ichiro Sakuma75,
  • Shu Takagi75,
  • Mathieu D. Santin76, 77,
  • Laurent Marsac78, 79,
  • Guillaume Maimbourg80,
  • Morgane Monfort77,
  • Benoit Larrat79,
  • Chantal François77,
  • Stéphane Lehéricy76, 77,
  • Mickael Tanter81,
  • Jean-Franҫois Aubry79,
  • Maria Eleni Karakatsani82,
  • Gesthimani Samiotaki82,
  • Shutao Wang82,
  • Camilo Acosta82,
  • Eliza R. Feinberg83,
  • Elisa E. Konofagou82, 84,
  • Zsofia I. Kovacs85,
  • Tsang-Wei Tu85,
  • Georgios Z. Papadakis85, 86,
  • William C. Reid86,
  • Dima A. Hammoud86,
  • Joseph A. Frank85, 87,
  • Zsofia i. Kovacs88,
  • Saejeong Kim88,
  • Neekita Jikaria88,
  • Michele Bresler88,
  • Farhan Qureshi88,
  • Joseph A. Frank88, 89,
  • Jingjing Xia91,
  • Po-Shiang Tsui90,
  • Hao-Li Liu91,
  • Juan C. Plata92,
  • Samuel Fielden93,
  • Bragi Sveinsson92,
  • Brian Hargreaves92,
  • Craig Meyer93,
  • Kim Butts Pauly92,
  • Juan C. Plata94,
  • Vasant A. Salgaonkar95,
  • Matthew Adams95,
  • Chris Diederich95,
  • Eugene Ozhinsky96,
  • Matthew D. Bucknor96,
  • Viola Rieke96,
  • Ari Partanen97, 98,
  • Andrew Mikhail98,
  • Lauren Severance98,
  • Ayele H. Negussie98,
  • Bradford Wood98,
  • Martijn de Greef100,
  • Gerald Schubert99,
  • Chrit Moonen100,
  • Mario Ries100,
  • Megan E. Poorman101, 102,
  • Mary Dockery101,
  • Vandiver Chaplin103, 102,
  • Stephanie O. Dudzinski101,
  • Ryan Spears101,
  • Charles Caskey104, 102,
  • Todd Giorgio101,
  • William Grissom101, 102,
  • Marcia M. Costa105,
  • Efthymia Papaevangelou105,
  • Anant Shah105,
  • Ian Rivens105,
  • Carol Box105,
  • Jeff Bamber105,
  • Gail ter Haar105,
  • Scott R. Burks106,
  • Matthew Nagle106,
  • Ben Nguyen106,
  • Michele Bresler106,
  • Joseph A. Frank106,
  • Scott R. Burks107,
  • Matthew Nagle107,
  • Ben Nguyen107,
  • Michele Bresler107,
  • Saejeong Kim107,
  • Blerta Milo107,
  • Joseph A. Frank107,
  • Nhan M. Le108,
  • Shaozhen Song110,
  • Kanheng Zhou108,
  • Ghulam Nabi109,
  • Zhihong Huang108,
  • Shmuel Ben-Ezra111,
  • Shani Rosen112,
  • Senay Mihcin113,
  • Jan Strehlow114,
  • Ioannis Karakitsios113,
  • Nhan Le113,
  • Michael Schwenke114,
  • Daniel Demedts114,
  • Paul Prentice113,
  • Sabrina Haase114,
  • Tobias Preusser114,
  • Andreas Melzer113,
  • Jean-Louis Mestas115,
  • Kamel Chettab118,
  • Gustavo Stadthagen Gomez116,
  • Charles Dumontet117,
  • Bettina Werle116,
  • Cyril Lafon115,
  • Fabrice Marquet119, 120,
  • Pierre Bour119, 120,
  • Fanny Vaillant119, 120,
  • Sana Amraoui119, 121,
  • Rémi Dubois119, 120,
  • Philippe Ritter119, 121,
  • Michel Haïssaguerre119, 121,
  • Mélèze Hocini119, 121,
  • Olivier Bernus119, 120,
  • Bruno Quesson119, 120,
  • Amit Livneh122,
  • Eitan Kimmel122,
  • Dan Adam122,
  • Justine Robin123,
  • Bastien Arnal124,
  • Mathias Fink123,
  • Mickael Tanter123,
  • Mathieu Pernot123,
  • Tatiana D. Khokhlova128,
  • George R. Schade125,
  • Yak-Nam Wang126,
  • Wayne Kreider126,
  • Julianna Simon126,
  • Frank Starr126,
  • Maria Karzova127,
  • Adam Maxwell125,
  • Michael R. Bailey126,
  • Vera Khokhlova126, 127,
  • Jonathan E. Lundt129,
  • Steven P. Allen129,
  • Jonathan R. Sukovich129,
  • Timothy Hall129,
  • Zhen Xu129,
  • George R. Schade131,
  • Yak-Nam Wang132,
  • Tatiana D. Khokhlova130,
  • Philip May131,
  • Daniel W. Lin131,
  • Michael R. Bailey132,
  • Vera Khokhlova133, 132,
  • Charlotte Constans135,
  • Thomas Deffieux134,
  • Mickael Tanter134,
  • Jean-Francois Aubry136,
  • Eun-Joo Park137,
  • Yun Deok Ahn137,
  • Soo Yeon Kang137,
  • Dong-Hyuk Park137,
  • Jae Young Lee137,
  • J. Vidal-Jove138, 141, 142,
  • E. Perich141,
  • A. Ruiz138,
  • A. Jaen139,
  • N. Eres141,
  • M. Alvarez del Castillo140,
  • Rachel Myers144, 143,
  • James Kwan143,
  • Christian Coviello143,
  • Cliff Rowe143,
  • Calum Crake144,
  • Sean Finn143,
  • Edward Jackson143,
  • Robert Carlisle144, 143,
  • Constantin Coussios144, 143,
  • Antonios Pouliopoulos145,
  • Caiqin Li145,
  • Marc Tinguely146,
  • Meng-Xing Tang145,
  • Valeria Garbin146,
  • James J. Choi145,
  • Paul C. Lyon147,
  • Christophoros Mannaris147,
  • Michael Gray147,
  • Lisa Folkes147,
  • Michael Stratford147,
  • Robert Carlisle147,
  • Feng Wu147,
  • Mark Middleton147,
  • Fergus Gleeson147,
  • Constantin Coussios147,
  • Sandra Nwokeoha148,
  • Robert Carlisle148,
  • Robin Cleveland148,
  • Yak-Nam Wang149,
  • Tatiana D. Khokhlova149,
  • Tong Li149,
  • Navid Farr149,
  • Samantha D’Andrea149,
  • Frank Starr149,
  • Kayla Gravelle149,
  • Hong Chen149,
  • Ari Partanen150,
  • Donghoon Lee149,
  • Joo Ha Hwang149,
  • Sophie Tardoski151,
  • Jacqueline Ngo151,
  • Evelyne Gineyts152,
  • Jean-Pau Roux152,
  • Philippe Clézardin152,
  • David Melodelima151,
  • Allegra Conti153, 154,
  • Rémi Magnin153, 155,
  • Matthieu Gerstenmayer153,
  • François Lux156,
  • Olivier Tillement156,
  • Sébastien Mériaux153,
  • Stefania Della Penna154,
  • Gian Luca Romani154,
  • Erik Dumont155,
  • Benoit Larrat153,
  • Tao Sun157, 158,
  • Chanikarn Power157,
  • Yong-Zhi Zhang157,
  • Jonathan Sutton157,
  • Eric Miller158,
  • Nathan McDannold157,
  • Oleg Sapozhnikov159, 160,
  • Sergey Tsysar159,
  • Petr V. Yuldashev159,
  • Vera Khokhlova159, 160,
  • Victor Svet161,
  • Wayne Kreider160,
  • Dongli Li162,
  • Antonio Pellegrino162,
  • Nik Petrinic162,
  • Clive Siviour162,
  • Antoine Jerusalem162,
  • Robin Cleveland162,
  • Peter V. Yuldashev163,
  • Maria Karzova163,
  • Bryan W. Cunitz164,
  • Barbrina Dunmire164,
  • Wayne Kreider164,
  • Oleg Sapozhnikov163, 164,
  • Michael R. Bailey164,
  • Vera Khokhlova163, 164,
  • Claude Inserra165,
  • Matthieu Guedra165,
  • Cyril Mauger166,
  • Bruno Gilles165,
  • Maxim Solovchuk167,
  • Tony W. H. Sheu168,
  • Marc Thiriet169,
  • Yufeng Zhou170,
  • Esra Neufeld171,
  • Christian Baumgartner171,
  • Davnah Payne171,
  • Adamos Kyriakou171,
  • Niels Kuster171, 172,
  • Xu Xiao173,
  • Helen McLeod174,
  • Andreas Melzer173,
  • Christopher Dillon175,
  • Viola Rieke176,
  • Pejman Ghanouni177,
  • Dennis L. Parker175,
  • Allison Payne175,
  • Vera A. Khokhova178, 179,
  • Peter V. Yuldashev178,
  • Ilya Sinilshchikov178,
  • Yulia Andriyakhina178,
  • Tatiana D. Khokhlova181,
  • Wayne Kreider179,
  • Adam Maxwell182,
  • Oleg Sapozhnikov178, 179,
  • Ari Partanen180,
  • Andrey Rybyanets183,
  • Natalia Shvetsova183,
  • Alex Berkovich184,
  • Igor Shvetsov183,
  • Oleg Sapozhnikov185,
  • Vera Khokhlova185,
  • Caroline J. Shaw186, 188,
  • Ian Rivens187,
  • John Civale187,
  • Dino Giussani186,
  • Gail ter Haar187,
  • Christoph Lees188,
  • Pierre Bour189, 190,
  • Fabrice Marquet189,
  • Valery Ozenne189,
  • Solenn Toupin189,
  • Bruno Quesson189,
  • Erik Dumont190,
  • Eugene Ozhinsky191,
  • Vasant Salgaonkar192,
  • Chris Diederich192,
  • Viola Rieke191,
  • Elena Kaye193,
  • Sebastien Monette193,
  • Majid Maybody193,
  • Govindarajan Srimathveeravalli193,
  • Stephen Solomon193,
  • Amitabh Gulati193,
  • Tobias Preusser194, 195,
  • Sabrina Haase194,
  • Mario Bezzi196,
  • Jürgen W. Jenne203,
  • Thomas Lango197,
  • Yoav Levy198,
  • Michael Müller199,
  • Giora Sat204,
  • Christine Tanner200,
  • Stephan Zangos201,
  • Matthias Günther194,
  • Andreas Melzer202,
  • Cyril Lafon205, 206,
  • Au Hoang Dinh205, 206,
  • Emilie Niaf207, 205,
  • Flavie Bratan209,
  • Nicolas Guillen208,
  • Rémi Souchon205, 206,
  • Carole Lartizien207,
  • Sebastien Crouzet209,
  • Olivier Rouviere205, 209,
  • Jean-Yves Chapelon205, 206,
  • Yang Han210,
  • Shutao Wang210,
  • Elisa E. Konofagou210, 211,
  • Thomas Payen212,
  • Carmine Palermo213,
  • Steve Sastra213,
  • Hong Chen212,
  • Yang Han212,
  • Kenneth Olive213,
  • Elisa E. Konofagou212, 214,
  • Johanna M. van Breugel216,
  • Martijn de Greef216,
  • Charles Mougenot217,
  • Maurice A. van den Bosch215,
  • Chrit Moonen216,
  • Mario Ries216,
  • Matthieu Gerstenmayer218,
  • Rémi Magnin218,
  • Benjamin Fellah218,
  • Denis Le Bihan218,
  • Benoit Larrat218,
  • Matthieu Gerstenmayer219,
  • Rémi Magnin219,
  • Sébastien Mériaux219,
  • Denis Le Bihan219,
  • Benoit Larrat219,
  • Steven P. Allen220,
  • Luis Hernandez-Garcia221, 220,
  • Charles A. Cain220,
  • Timothy Hall220,
  • Erasmia Lyka222,
  • Delphine Elbes222,
  • Christian Coviello222,
  • Robin Cleveland222,
  • Constantin Coussios222,
  • Kanheng Zhou223,
  • Nhan M. Le223,
  • Chunhui Li223,
  • Zhihong Huang223,
  • Satoshi Tamano224, 225,
  • Hayato Jimbo224,
  • Takashi Azuma226,
  • Shin Yoshizawa224,
  • Keisuke Fujiwara225,
  • Kazunori Itani225,
  • Shin-ichiro Umemura227,
  • Christakis Damianou228,
  • Marinos Yiannakou228,
  • Nicholas Ellens229,
  • Ari Partanen229, 230,
  • Dan Stoianovici231,
  • Keyvan Farahani229, 232,
  • Zulfadhli Zaini233,
  • Ryo Takagi234,
  • Shin Yoshizawa234,
  • Shin-ichiro Umemura233,
  • Shenyan Zong235,
  • Guofeng Shen236,
  • Ron Watkins237,
  • Aurea Pascal-Tenorio238, 246,
  • Matthew Adams239, 247,
  • Juan C. Plata240,
  • Vasant Salgaonkar241,
  • Peter Jones242,
  • Kim Butts-Pauly243,
  • Chris Diederich244,
  • Donna Bouley245,
  • Andrey Rybyanets248,
  • Guoxin Ren249,
  • Wei Guo249,
  • Guofeng Shen249,
  • Yazhu Chen249,
  • Chung-Yin Lin250,
  • Han-Yi Hsieh250,
  • Kuo-Chen Wei250,
  • Hao-Li Liu250,
  • Camille Garnier251, 253,
  • Gilles Renault252,
  • Navid Farr254,
  • Ari Partanen255,
  • Ayele H. Negussie254,
  • Andrew Mikhail254,
  • Reza Seifabadi254,
  • Emmanuel Wilson256,
  • Avinash Eranki256,
  • Peter Kim256,
  • Bradford Wood254,
  • Dennis Lübke257,
  • Jürgen W. Jenne257, 258,
  • Peter Huber258,
  • Matthias Günther257,
  • Dennis Lübke259,
  • Joachim Georgii259,
  • Michael Schwenke259,
  • Caroline V. Dresky259,
  • Julian Haller260,
  • Matthias Günther259,
  • Tobias Preusser259,
  • Jürgen W. Jenne259,
  • Avinash Eranki261, 262,
  • Navid Farr262,
  • Ari Partanen263,
  • Pavel Yarmolenko261,
  • Ayele H. Negussie262,
  • Karun Sharma261,
  • Haydar Celik262,
  • Bradford Wood262,
  • Peter Kim261,
  • Guofeng Li264,
  • Weibao Qiu264,
  • Hairong Zheng264,
  • Meng-Yen Tsai265,
  • Po-Chun Chu265,
  • Hao-Li Liu265,
  • Taylor Webb266,
  • Urvi Vyas267,
  • Kim Butts Pauly266, 267,
  • Matthew Walker268, 269,
  • Jidan Zhong269,
  • Thomas Looi270,
  • Adam C. Waspe270,
  • James Drake270,
  • Mojgan Hodaie269, 268,
  • Feng-Yi Yang271,
  • Sin-Luo Huang271,
  • Yuval Zur272,
  • Alexander Volovick273,
  • Benny Assif273,
  • Christian Aurup274,
  • Hermes Kamimura275, 274,
  • Shutao Wang274,
  • Hong Chen274,
  • Camilo Acosta274,
  • Antonio A. Carneiro275,
  • Elisa E. Konofagou274,
  • Alexander Volovick276,
  • Javier Grinfeld276,
  • David Castel277,
  • Sven Rothlübbers278, 279,
  • Julia Schwaab278,
  • Christine Tanner280,
  • Senay Mihcin281,
  • Graeme Houston281,
  • Matthias Günther279, 278,
  • Jürgen W. Jenne279, 278,
  • Eugene Ozhinsky282,
  • Matthew D. Bucknor282,
  • Viola Rieke282,
  • Haim Azhari283,
  • Noam Weiss283,
  • Jacob Sosna284,
  • S. Nahum Goldberg284, 285,
  • Victor Barrere286, 287,
  • David Melodelima286, 287,
  • Kee W. Jang288,
  • Scott R. Burks288,
  • Zsofia I. Kovacs288,
  • Tsang-Wei Tu288,
  • Bobbi Lewis288,
  • Saejeong Kim288,
  • Matthew Nagle288,
  • Neekita Jikaria288,
  • Joseph A. Frank288,
  • Yufeng Zhou289,
  • Xiaotong Wang289,
  • Yun Deok Ahn290,
  • Eun-Joo Park290,
  • Dong-Hyuk Park290,
  • Soo Yeon Kang290,
  • Jae Young Lee290,
  • Visa Suomi291,
  • Elisa E. Konofagou292,
  • David Edwards291,
  • Robin Cleveland291,
  • Zahary Larrabee293,
  • Matthew Eames293, 294,
  • Arik Hananel293,
  • Jean-Franҫois Aubry295, 296,
  • Boaz Rafaely297,
  • Alexander Volovick298,
  • Javier Grinfeld298,
  • Eitan Kimmel299,
  • Rasha Elaimy Debbiny299,
  • Carmel Zeltser Dekel300,
  • Michael Assa300,
  • Eitan Kimmel300,
  • George Menikou301,
  • Christakis Damianou302,
  • Petros Mouratidis303,
  • Ian Rivens303,
  • Gail ter Haar303,
  • José A. Pineda-Pardo304,
  • Marta Del Álamo de Pedro304,
  • Raul Martinez304,
  • Frida Hernandez304,
  • Silvia Casas305,
  • Carlos Oliver305,
  • Patricia Pastor305,
  • Lidia Vela304,
  • Jose Obeso304,
  • Paul Greillier306,
  • Ali Zorgani306,
  • Rémi Souchon306,
  • David Melodelima306,
  • Stefan Catheline306,
  • Cyril Lafon306,
  • Vyacheslav Solovov307,
  • Michael O. Vozdvizhenskiy307,
  • Andrew E. Orlov307,
  • Chueh-Hung Wu308,
  • Ming-Kuan Sun308, 310,
  • Tiffany T. Shih309,
  • Wen-Shiang Chen308, 310,
  • Fabrice Prieur311, 312,
  • Arnaud Pillon313,
  • Jean-Louis Mestas312, 314,
  • Valerie Cartron313,
  • Patrick Cebe313,
  • Nathalie Chansard313,
  • Maxime Lafond312,
  • Cyril Lafon312, 314,
  • Claude Inserra315,
  • Pauline Muleki Seya316,
  • Wen-Shiang Chen317,
  • Jean-Christophe Bera315,
  • Tanguy Boissenot318,
  • Benoit Larrat319,
  • Elias Fattal318,
  • Alexandre Bordat318,
  • Helene Chacun318,
  • Claire Guetin318,
  • Nicolas Tsapis318,
  • Kazuo Maruyama320,
  • Johan Unga320,
  • Ryo Suzuki320,
  • Cécile Fant321,
  • Maxime Lafond321,
  • Bernadette Rogez321,
  • Jacqueline Ngo321,
  • Cyril Lafon321,
  • Jean-Louis Mestas321,
  • Mercy Afadzi322,
  • Ola Finneng Myhre323,
  • Siri Vea323,
  • Astrid Bjørkøy322,
  • Petros Tesfamichael Yemane322,
  • Annemieke van Wamel322,
  • Sigrid Berg324, 323,
  • Rune Hansen324, 323,
  • Bjørn Angelsen323 and
  • Catharina Davies322
Journal of Therapeutic Ultrasound20175(Suppl 1):15

https://doi.org/10.1186/s40349-016-0079-2

Published: 17 March 2017

ORAL PRESENTATIONS

O1 The role of bubbles and cavitation in therapy ultrasound

Brian Fowlkes

Basic Radiological Sciences Division, Department of Radiology, University of Michigan, Ann Arbor, Michigan, USA

When exposed to sufficiently high ultrasound pressures, microbubbles can be generated spontaneously in tissue and undergo inertial cavitation where collapses result in physical effects. These effects range from petechial haemorrhage to complete cellular disruption, termed Histotripsy, depending on ultrasound parameters. This presentation will explore the mechanisms associated with histotripsy along with the tissue effects and the wide range of potential applications for this mechanical disruption method.

O2 Challenges for clinical trials in therapeutic ultrasound, the need for an evidence base, & trial design

Pejman Ghanouni

Stanford School of Medicine, Stanford, California, USA

In this lecture, we will compare the design of clinical trials that led to approval of MR guided focused ultrasound for the treatment of uterine fibroids and osseous metastases. The impact of these trials on the evidence base, and thus on adoption by users and coverage by insurers will be compared. We will also review the process of expanding approved FUS applications, either via investigator- or industry-initiated studies or through off-label clinical use.

O3 Prostate HIFU – current status

Narendra Sanghvi

SonaCare Medical, Indianapolis, Indiana, USA

In this lecture, present status of focused ultrasound for the treatment of localized prostate cancer ablation will be discussed. High intensity focused ultrasound (HIFU) has been used for the ablation of prostate over two decades and has treated thousands of prostate cancer patients. Meanwhile, prostate cancer management is undergoing significant improvements as molecular markers, targeted biopsy and advanced multi-parametric MRI are routinely used to accurately localize the prostate cancer. These advances offer a unique opportunity for focal ablation of localized prostate cancer with HIFU as it plays a significant role in reducing morbidity and treatment cost. This presentation will focus on hardware design, software architecture and HIFU features of the devices. Presentation will demonstrate localization of prostate with ultrasound imaging, treatment planning with 3D volumetric rendering of the prostate with ultrasound and MRI fusion techniques for focal treatment and finally HIFU dose setting with guidance using real time Tissue Change Monitoring (TCM) with the Sonablate device. The presenter will encourage exchange of ideas and discussion for research topics.

O4 Enhancement of drug delivery - clinical challenges and solutions

Constantin Coussios1, Paul C. Lyon1, Michael Gray1, Christophoros Mannaris1, Marie de Saint Victor1, Eleanor Stride1, Robin Cleveland1, Robert Carlisle1, Feng Wu4, Mark Middleton3, Fergus Gleeson2

1Institute of Biomedical Engineering, University of Oxford, Oxford, United Kingdom; 2Department of Radiology, Churchill Hospital, Oxford, United Kingdom; 3Department of Oncology, University of Oxford, Oxford, United Kingdom; 4HIFU Unit, Churchill Hospital, Oxford, UK

There are four key clinical challenges in optimizing drug based strategies: (i) achieving prolonged blood circulation of the therapeutic to enable active or passive accumulation in the target tissue; (ii) mediating triggered release or activation, or active accumulation of the therapeutic, to maximize its concentration at the target site whilst reducing off target side effects; (iii) enabling successful transport of the therapeutic from the blood stream into the target tissue, and achieving a homogenous distribution in that target tissue and (iv) where necessary, further enabling penetration of the therapeutic into the cell. The potential of therapeutic ultrasound with, or without, sonosensitive microparticle or nanoparticle formulations to address these challenges will be explored.

O5 Neuromodulation with ultrasound for beginners

Jean-Franҫois Aubry

Institut Langevin, Paris, France

In this lecture, the use of transcranial low intensity focused ultrasound for neuromodulation will be discussed. A historical review will be presented, with an emphasis on the experimental setups and the acoustical parameters. Models ranging from slice cultures to intact rodents and primates will be presented, together with recent trials on humans. Potential mechanisms will be described. Based on our own experience, exciting successful neuromodulation as well as disappointing failures will be presented.

O6 Thermometry in ultrasound fields, challenges & solutions in vivo, ex vivo & everywhere else!

Kim Butts Pauly

Stanford School of Medicine, Stanford, California, USA

This talk will cover image-based thermometry methods for guiding focused ultrasound. MR Thermometry based on the proton resonance frequency shift with temperature is linear and reversible in aqueous tissues and is utilized in clinical practice with common temperature resolutions of 1°C. Ultrasound based thermometry based on the speed of sound change with temperature can be used to 45°C in aqueous tissue. Both methods are sensitive to the presence of fat within the aqueous tissue, as well as motion. This talk will cover these basic concepts as well as their use in clinical practice.

O7 Motion compensation

Chrit Moonen

Center for Imaging Sciences, Imaging Division, University Medical Center, Utrecht, the Netherlands

Motion leads to several challenges for HIFU treatment. In this lecture, the effects of respiratory, cardiac, and peristalsis related motion on MRI thermometry will be discussed. In addition, methods for tracking the moving target with the HIFU will be described, as well as gating strategies.

O8 MRgFUS crossing the chasm from proof of concept to mainstream treatment alternative

Jacob Vortman

InSighTec, Haifa, Israel

MRgFUS treatment is a disruptive, non-invasive, outpatient treatment alternative that is capable to treat tumours and functional disorders under real-time monitoring and control via MR thermometry.

Transforming this breakthrough technology from the lab to a mainstream treatment alternative requires gaining the support and the agreement to change by a whole array of stakeholders in different areas some interrelated and some conflicting.

The stakeholders’ current position, changes they will need to go through and possible changes engines are mentioned below:

▪ Physicians (surgeons) need to transform into image guided surgeons where knowledge and understanding of the disease play a dominant role in the procedure outcome. The benefit is the confidence in the safety and efficacy coupled with income that wouldn’t decline. The change engine could be the patients and the payers.

▪ Payers should benefit from covering MRgFUS by saving cost and addressing patients’ demands. The change engine in this case should be the patients, governments and physicians.

▪ Governments should see the benefit of very fast recovery, very low level of adverse events and productivity enhancement. In this case patients and physicians should drive the change.

▪ Patients should benefit from safer treatment, fast recovery next day back to your life, minimal trauma and morbidity. The significant benefit to them should transform them to the dominant driver of this change. They will need to influence physicians, payers and providers to adopt this new treatment.

▪ Providers should adopt the technology and provide this treatment since the data exist proving safety and efficacy, proven cost savings and physicians and patients demand.

The current Medical ecosystem is biased against new technologies since the incumbent system/treatments are reimbursed while the new technologies are not. Could governments perform economic analysis and if found beneficial (example: saving money and improving productivity) decide on limited 2 years reimbursement during which RCT data will be collected based on which private insurance will decide to cover. This model should incentivize the physician, payers and providers to try the new technology.
Fig. 1 (abstract O8).

Overcoming the resistant to change - is there a strategy that could bring all the different stakeholders to combine and align efforts?

O9 MR guided focused ultrasound treatment of soft tissue tumours of the extremities

Pejman Ghanouni

Stanford School of Medicine, Stanford, California, USA

In this lecture, the use of MR guided focused ultrasound for the treatment of soft tissue and osseous tumours will be discussed. Results of treatment of these different types of tumours will be reviewed, including lessons learned from challenging patient treatments. Technical aspects of all parts of a treatment, including patient preparation, positioning, imaging, planning, thermometry, and methods of evaluation, will be described. The talk will also focus on methods developed to address these current challenges and opportunities for future development.

O10 Non-invasive, non-destructive FUS-induced neuro-modulation assessed by recording auditory evoked potentials – initial experience in small/large animals

Shirley Sharabi1,2, Dianne Daniels1,2, David Last1, David Guez1, Yoav Levy3, Alexander Volovick3, Javier Grinfeld3, Itay Rachmilevich3, Talia Amar3, Zion Zibly1, Yael Mardor1,2, Sagi Harnof1,2

1Sheba Medical Center, Ramat-Gan, Israel; 2Tel-Aviv University, Tel-Aviv, Israel, 3InSighTec, Haifa, Israel

Objectives

MR guided focused ultrasound (MRgFUS) has been extensively studied in recent years as a non-invasive treatment modality. Initial clinical trials have indicated promising treatment response to ablative FUS treatments of patients with brain tumours, neuropathic pain, essential tremor, obsessive compulsive disorder, and Parkinson’s disease. Apart from the ablative applications of FUS, this technology has been extensively evaluated for less destructive applications such as thrombolysis, blood–brain barrier disruption for increased drug delivery, and recently also neuro-modulation.

The objective of the presented study was to demonstrate non-invasive, non-destructive, reversible FUS-induced neuro-modulation in small (rats) and large (pigs) animals by inducing temporary suppression of auditory evoked potentials.

Methods

All animal experiments were performed under full anaesthesia. Rats were anesthetized by Xylazine/Ktalar and pigs were anesthetized by Propofol. Rat’s heads were shaved prior to treatment but they did not undergo craniotomy. Pigs underwent craniotomy to avoid FUS reflection/aberration by the skull. EEG was recorded using 3 small cup-shaped electrodes attached to the skin/dura with metal particles-containing gel for optimal sound conductivity (Fig. 2).

The audio stimulation system consisted of a pulse generator connected to the EEG trigger input and to speakers placed near the animal ears, producing a square-wave form at 10 KHz, resulting in 150 “click” sounds per min (Fig. 2). Each measurement consisted of 200 repetitions enabling acquisition of a full measurement in 1’20” min.

The ExAblate Neuro system (InSightec, Tel Aviv, Israel) is a combination of a standard MRI scanner and a FUS delivery system. The FUS device is in the shape of a helmet consisting of 1024 transducers which deliver US energy in the form of “sonications”. The system is designed to provide real-time therapy, planning, thermal dosimetry, and closed loop therapy control. Treatment starts with conventional MRI scans, displayed on the ExAblate computer, used to determine regions of interest of the target volume. During the procedure, the beam path is periodically reviewed to confirm the planned direction through the tissue. The set of sonication volumes is sequentially applied to cover the entire planned volume. The current experiments were performed with a modified ExAblate version developed for neuro-modulation as part of the MAGNET programs supported by the Israeli Ministry of Commerce.

Baseline auditory evoked potentials were recorded by EEG prior to FUS treatment, with the animals in the prone position. The animals were then placed in the supine position, with the skull dipped in degassed water at the centre of the FUS system, for localization MRI scanning followed by FUS treatment. The animals were then returned to the prone position for continuous post-treatment EEG recordings. Rats which did not show recovery of the auditory evoked potentials 30–60 min post treatment were monitored again 48 hours or 1 week post treatment.

The animals were treated by FUS for 52 sec using the Exablate Neuro system at 220 KHz, 12 W, and 100 ms on/ 2900 ms off pulses. Two rats were treated in the thalamus region (targeted at deep auditory tracks) and another two in the frontal cortex region (targeted at peripheral auditory tracks). Two sham rats underwent a similar procedure without activation of the FUS system. One pig was treated in the thalamus region and another in the right motor cortex region.

Results

Auditory evoked potential EEG signals shapes varied from one animal to the other but all were detected 2–10 ms after the trigger. The maximal peak-to-peak height was calculated for each measurement.

The sham rats showed no significant change in the auditory evoked potentials EEG signal.

The rats treated in the thalamus regions showed 50% and 65% suppression of the baseline auditory evoked potentials EEG signal. The first showed no recovery 2 hours post treatment with full recovery measured 1 week post treatment. The second showed no recovery for 1 hour post treatment and full recovery 48 hours post treatment. The rats treated in the cortex regions showed 50% and 67% suppression of the baseline EEG signal. The first showed no recovery 30 min post treatment with full recovery measured 1 week post treatment. The second showed initial recovery 14 min post treatment reaching full recovery within 28 min post treatment.

The pig treated in the thalamus region showed 90% suppression of the baseline signal with no recovery 30 min post treatment. The second pig, treated in the cortex region, showed complete suppression of the baseline signal immediately post treatment with initial recovery noted 18 min post treatment, reaching full recovery 63 min post treatment (Fig. 3).

Conclusions

Our preliminary results suggest that reversible neuro-modulation by non-invasive FUS is feasible. Full recovery was noted in all 4 treated rats and in 1 of the 2 treated pigs. Unfortunately we were not able to monitor the first pig for more than 30 min post treatment.
Fig. 2 (abstract O10).

See text for description

Fig. 3 (abstract O10).

See text for description

O11 Biophysical dissection of ultrasonic neuromodulation mechanisms

Michael Plaksin, Yoni Weissler, Shy Shoham, Eitan Kimmel

Faculty of Biomedical Engineering & Russell Berrie Nanotechnology Institute, Technion – Israel Institute of Technology, Haifa, Israel

Objectives

Low intensity US can noninvasively suppress or excite central nervous system (CNS) activity using different combinations of stimulation parameters. While applications are already emerging, the underlying biophysics remains unclear regarding the relative contribution of possible mechanisms: extracellular bubble cavitation, thermal effects, acoustic radiation pressure and US-induced intramembrane cavitation within the bilayer membrane (the bilayer sonophore or BLS model). Interestingly, both radiation pressure and intramembrane cavitation can induce plasma membrane capacitance changes. Here, we use detailed predictive modelling and find that only intramembrane cavitation can explain all the observed aspects of ultrasonic neuromodulation.

Methods

We analyzed the relevant experimental literature using modified Rayleigh–Plesset intramembrane cavitation BLS biomechanics and acoustic radiation pressure gradients (RPG) - induced membrane dynamics. By coupling these biomechanical models to biophysical membrane models we predict dynamical biophysical responses of artificial bilayer membranes, and of three common neocortical single cell Hodgkin-Huxley type models: i) Regular Spiking (RS) cortical pyramidal neuron, ii) Fast Spiking (FS) cortical inhibitory neuron and iii) Low Threshold Spiking (LTS) cortical inhibitory neuron, RS-FS-LTS Hodgkin-Huxley based network model and CNS axon model. In addition, live brain tissue RPG subjected areal strains were evaluated in a viscoelastic brain model.

Results

Only the Neuronal Intramembrane Cavitation Excitation (NICE) models were able to explain US-induced action potential generation through BLS-type pulsating nano-bubbles inside the bilayer plasma membrane: the leaflets' periodic vibrations induce US-frequency membrane capacitance and potential oscillations, leading to slow charge accumulation across the membrane (on a time scale of tens of milliseconds), until action potentials are generated. In contrast, the analysis of RPG-induced membrane capacitance variations associated with membrane area changes explain artificial membrane results, but were found to be highly unlikely sources for neural excitation, when considering the areal strains expected to form in brain tissue during normal sonication. Further, the NICE-LTS inhibitory neurons show a much higher relative sensitivity to sparse ultrasonic stimulation compared to the other neurons, resulting from their T-type voltage gated calcium channels. This model-based prediction was found to explain the results of a significant body of suppression and excitation experimental studies, including in humans.

Conclusions

These results provide a unified theoretical framework for a large body of experiments in multiple preparations across the field of US neuromodulation, lending further support to the hypothesis that intramembrane cavitation is responsible for ultrasonic neuromodulation. They could thus pave the way towards new CNS therapeutic protocols, using the only method that currently allows targeted non-invasive neuromodulation with millimetre spatial resolution essentially anywhere in the brain.

O12 Ultrasonic stimulation of mammalian retina in-vitro

Omer Naor1,2, Nairouz Farah3, Shy Shoham2

1ELSC Center for Brain Sciences, Hebrew University, Jerusalem, Israel; 2Biomedical Engineering, Technion- Israel Institute of Technology, Haifa, Israel; 3Faculty of Life Sciences, Bar Ilan University, Ramat Gan, Israel

Objectives

Following previous in vivo stimulation of the retina, we aimed to achieve a first direct measurement of the response of mammalian retinal neurons to ultrasonic (US) stimuli, and to study and characterize this response.

Methods

We coupled a high-density phased array (986 elements on a 25x35 mm2 area) to a system for multi-electrode-array (MEA) recording with 256 contacts. Mouse retinas were dissected and placed on the MEA, and sonicated at 2.3 MHz, applying varying durations and intensities, as well as stimulated by light. The acquired data were processed to detect action potentials (spikes) elicited by retinal ganglion cells, and analysed to reveal the relations between the stimuli and the responses.

Results

We found prominent spike responses for stimuli in the range of 4.3-7.3 W/cm2 and 0.5-1 s, which disappeared when the focus was steered 1.5 mm away. Furthermore, we found that the relation between the response strength and the stimulation intensity, or duration, followed a logistic sigmoid curve, while the response latency was described by a decreasing exponent. Lastly, we found indications that the observed responses to US stimuli are related to the "2nd OFF" component in the responses to light stimuli.

Conclusions

These findings are the first direct demonstration of the response of the mammalian retina to US stimulation. The properties of the US transducer and the stimulation frequency indicate that non-invasive US stimulation of human retina is feasible, and may potentially evolve as an important tool for diagnosis and treatment of retinal diseases.

O13 Motor response elicitation and pupil dilation using megahertz-range focused ultrasound neuromodulation

Christian Aurup1, Hermes Kamimura2,1, Shutao Wang1, Hong Chen1, Camilo Acosta1, Antonio A. Carneiro2, Elisa E. Konofagou1

1Columbia University, New York, New York, USA; 2Universidade de São Paulo, São Paulo, Brasil

Objectives

Using transcranial focused ultrasound for the modulation of brain activity has been identified as a possible non-invasive means of treating neurological disorders. Most studies involving sedate rodents use frequencies in the kilohertz range, which allow for optimal transmission of acoustic power through the skull. The trade-off of using lower frequencies involves a lack of target specificity. Higher frequencies must be used in order to modulate activity in a more highly-specified manner. This study demonstrates that focused ultrasound in the megahertz range can be used to evoke motor- and cognitive-related responses in mice under deep anaesthesia by targeting specific brain structures. Contralateral-paired hind limb movements were observed when stimulating cortical regions, demonstrating the ability of MHz-range FUS to stimulate activity in highly-localized brain regions. Additionally, pupil dilation was observed when deep-seated anxiety-related structures were targeted, demonstrating the ability of FUS to modulate cognitive activity in a highly-specified manner.

Methods

For this study, wild-type adult male mice were anesthetized with intraperitoneal injections of sodium pentobarbital (65 mg/kg) and fixed in a stereotaxic frame. A single-element FUS transducer with fundamental frequency of 1.94 MHz was fixed to a 3D positioning system for accurate navigation through the brain. A 6x6 mm grid centred +2 mm rostral of the lambda skull suture was sonicated in a random order using a centre frequency of 1.9 MHz, pulse repetition frequency of 1 kHz, 50% duty cycle, 1 second pulse duration, 1 second inter-pulse interval for a total of 10 pulse repetitions. The acoustic pressure applied was varied in order to evaluate thresholds for eliciting physiological responses like motor movement, eye movement, or pupil dilation. Motor movements were validated using video recordings and electromyography via needle electrodes implanted into the biceps femoris of both hind limbs. Videos were recorded using a high-resolution camera focused at the right eye and processed to measure eye movements or changes in pupil size.

Results

The minimum acoustic pressure required to elicit motor movements was 1.45 MPa when targeting the somatosensory cortex, calibrated using an excised mouse skull. Higher pressures increased the success rate from 20% (at the 1.45 MPa threshold) to 70% (1.79 MPa). Targeting eye-motor and anxiety related regions of the brain elicited eye movements and pupil dilations up to 20%. Sonicating the superior colliculus resulted in both eye movement and pupil dilation at a lower threshold pressure (1.20 MPa) than the hippocampus and locus coeruleus which required pressures greater than 1.80 MPa.

Conclusions

This study successfully demonstrated that MHz-range transcranial focused ultrasound can be used to elicit motor- and cognitive-related physiological responses with high specificity in mice in vivo. It was also shown that the success rate of stimulation increased with acoustic pressure for motor movements associated with cortical activity modulation but highly depends on the region of the brain targeted. These findings emphasize the complex and yet to be determined mechanism of action involved in ultrasonic neuromodulation.
Fig. 4 (abstract O13).

Evaluation of the pressure threshold and success rate associated with applying FUS to location within the somatosensory cortex. This location resulted in contralateral hind-limb movement relative to the sonication site. Moving the transducer symmetrically about the midline resulted again in contralateral movement relative to the new sonication site

Fig. 5 (abstract O13).

Superior colliculus (top) threshold determined to be approximately 1.2 MPa while the locus coeruleus (bottom) was evaluated to be greater than 1.8 MPa

O14 Thermal dose effects by MR-guided focused ultrasound on the pig brain tissue - preliminary results

Dong-Guk Paeng1,3, Zhiyuan Xu2, John Snell1, Anders H. Quigg1, Matthew Eames1, Changzhu Jin3, Ashli C. Everstine4, Jason P. Sheehan2, Beatriz S. Lopes5, Neal Kassell1

1Focused Ultrasound Foundation, Charlottesville, Virginia, USA; 2Neurosurgery, University of Virginia, Charlottesville, Virginia, USA; 3Ocean System Engineering, Jeju National University, Jeju, Korea (the Republic of); 4Biology, University of Virginia, Charlottesville, Virginia, USA; 5Pathology, University of Virginia, Charlottesville, Virginia, USA

Objectives

The objective of this research is to investigate the effects of thermal dose (TD) delivered by magnetic resonance-guided focused ultrasound (MRgFUS) on in vivo pig brain tissue. In current clinical applications of transcranial MRgFUS systems, continuous acoustic wave emission is used to heat brain tissue to peak temperatures over 58°C. However, there are some situations where it has proven difficult to reach the desired peak temperature due to high absorption of acoustic energy by skull bone. There are reports that thermal effects on tissue are well correlated with thermal dose, which suggest that treatment delivery could be prescribed in terms of thermal dose rather than peak temperature or electric/acoustic power. It is also been demonstrated that the thermal dose threshold for permanent tissue damage is about 240 cumulative equivalent minutes (CEM) at 43°C for most of tissue. Currently available transcranial MRgFUS systems only allow the prescription of acoustic power and duration. In order to investigate the effects of thermal dose on in vivo brain tissue, we have developed a closed-loop control system to allow prescription thermal dose. This system monitors tissue heating via MR thermometry and provides pulse width modulation of output acoustic power in order to hold target tissue at a fixed temperature, and hence receives a nearly constant dose rate.

Methods

A FUS system (ExAblate 4000 Neuro 650 kHz system, InSightec) was used for sonication and an MRI system (Discovery MR75-3.0T, GE Medical systems) was used for thermometry and pre- and post-imaging. A closed-loop control system was implemented on a personal computer to control pulse width modulation of the FUS system acoustic power in order to maintain a specified temperature based on the MR thermometry. Accumulated thermal dose was calculated in real time and used to stop the sonication so that a prescribed thermal dose was delivered to the targeted tissue. Phantom studies were performed to test the control system to prepare for animal experiments. One acute and six chronic experiments (with three day survival) were conducted to observe the effects of TD on pig brain by behaviour observation and post MR imaging of the brain (1 hour and 70 hours post procedure). Craniotomy was performed to create an acoustic access window, and sonication was applied on 4 spots in the thalamus of each pig. Histology was also performed to compare it with MR imagery. Temperature in the pig brain tissue was estimated by rectal temperature for the MR thermometry baseline. TD was varied from 7 to 200 CEM while the target temperature was changed from 46 to 52 °C with appropriate acoustic power depending on target position and individual pig. This study was approved by the University of Virginia Institutional Animal Care and Use Committee.

Results

From the acute experiment, we could observe the lesions on MR images after 1 hour of sonication and histology subsequently confirmed the lesions. For the chronic experiments, no obvious problem was observed in the behavior of any of the six animals. Eighteen sonication spots in 5 pigs were analyzed through MR images. One pig experiment failed to control temperature due to introduction of air bubbles between the brain and scalp during surgery procedure, and 2 sonication spots were excluded due to technical problems. Large tissue changes were observed in MR images in all 6 spots over 100 CEM.

The diameter of those tissue changes in MR T2-weighted axial images were measured and averaged to 2.9 ± 0. 4 mm. There is inconsistency in generating lesions for TD below 100 CEM. No lesion was shown in some lower TD from 7 CEM and 61 CEM, while some smaller lesions (<2 mm in lesion diameter) were shown in TD from 18 CEM to 85 CEM except one large tissue change of 3.5 mm in diameter at 31 CEM. Some tissue changes were shown in both post MR images after 1 hour and 70 hours of sonication, while some were visible only at the 70 hour time point. Histology of 3 pig experiments is now available and the histology reports support the tissue changes and lesions in MR images. Lesion diameters in MR T2-weighted axial images versus TD in CEM are shown in Fig. 6 for all the results from the chronic pig study.

Conclusions

These preliminary results from pig brain tissue generally confirmed the previous results from rabbit brain tissue in generating tissue changes over a certain TD, even though there are some differences in the FUS systems and the experimental procedures and analysis. For lower thermal dose below 61 CEM, there is significant variability in generating of tissue changes, while large tissue changes whose average diameter is 2.9 mm were observed in MR T2-weighted axial images for higher TD over 100 CEM, which were reported with similar tendency but a little difference in TD from the rabbit brain study. These results may contribute to open the way to prescribe the thermal dose rather than peak temperature or acoustic power for brain treatments, and expand the treatment envelope beyond the current limitations in selecting targets and patients. This project is ongoing and will be further pursued with additional experiments for consolidation of the results and analysis.
Fig. 6 (abstract O14).

Pig chronic study results showing the relations of lesion diameter (mm) based on enhancing region appearing in T2-weighted axial MR images with applied thermal dose in CEM.

O15 In vivo feasibility study of boiling histotripsy with clinical Sonalleve system in a neurological porcine model

Thomas Looi1,4, Vera Khokhlova2, Charles Mougenot5, Kullervo Hynynen3, James Drake1

1Hospital for Sick Children, Toronto, Ontario, Canada; 2University of Washington, Seattle, Washington, USA; 3Sunnybrook Research Institute, Toronto, Ontario, Canada; 4University of Toronto, Toronto, Ontario, Canada; 5Philips Healthcare, Markham, Ontario, Canada

Objectives

To determine if a clinical focused ultrasound system (Philips Sonalleve) can be used to perform mechanical liquefaction of brain tissue for neurological lesioning through a simulated fontanelle in a porcine model (simulating neo-natal patients). This work will determine the power of the system required to induce lesions using a boiling histotripsy (BH) pulsing protocol. Post-treatment, the lesion volume and border will be measured with MRI imaging and histological examination.

Methods

A porcine model was used as the in vivo model with a maximum weight < 6.8 kg (4.9 - 6.8kg). A horse-shoe incision and blunt dissection was used to expose the skull. A craniotomy was performed to create a 4–5 cm2 opening in the skull simulating the fontanelle in a neonatal patient. A degassed mixture of ultrasound gel and water (ratio 10:1) was poured on top of the dura to ensure good acoustic coupling. The scalp was sutured closed with 2–0 Vicryl cutting needle. The animal was placed supine feet first with the craniotomy centred about the Sonalleve V2 system with Flex-M surface coils. Pre-treatment T1-weighted (T1-w), T2-weighted (T2-w) and T2*-weighted (T2*-w) MRI imaging was conducted as a baseline. Each animal was treated at four cluster locations where each cluster consisted of seven sonication points; one point in the centre and six points uniformly distributed over a 4-mm diameter circle. The clusters were located approximately 15 mm deep in the brain, 7 mm off the midline, and separated by 14 mm in a rectangular geometry. In initial treatment on the first animal, the power was increased from 100 to 500 W for each cluster. After initial analysis, the treatment was repeated on second animal with refined power levels of 325, 350, 375, and 425 W. The treatment sequence consisted of 12000 pulses of 1.2 MHz frequency, 1 and 10 ms pulse duration, and 1% duty cycle for both 1 and 10 ms pulse duration. These protocols have been shown to generate BH lesions in ex vivo bovine liver in another Sonalleve system. During treatment, MR thermometry was used to monitor for surface, focal, and far field heating. A dedicated MATLAB-based interface was connected to the Sonalleve cavitation sensor to detect the signal generated during treatment points. After treatment, post T1-w, T2-w and T2*-w MRI scans were completed for comparison. The animals were euthanized, perfusion fixated and their brains were removed for histology. The brain specimens were cut at the centre for the treatment clusters to get a cross-sectional coronal view where each slice was 5 microns. The slides were stained with haematoxylin and eosin (H&E) and examined for lesion presence, blood and border definition.

Results

A total of 4 piglets were sonicated with the following configuration: 1 piglet with 5 clusters (100 – 500 W), 2 piglets with 4 clusters (325-425 W) using 10-ms long pulses and 1 piglet with 4 clusters and 1-ms long pulses. For all power levels, the MR-measured temperature in the near or far field of the treatment was below the noise level. For 10-ms long pulses, 100 and 200 W acoustic powers, no noticeable imaging change was observed during sonication and post-treatment MR evaluation. As power was increased from 300 to 400 W, a temperature increase of up to 5C was measured at the focus. With more discrete power levels, it appeared at 375W that the lesion was more contained whereas higher power levels created wider areas of tissue change. It was observed the timing of MR magnitude of the target cluster changed as the power level increased where the tissue change occurred at 15 s at 300 W, 12 s at 350 W, 10 s at 375 W and 5 s at 425 W. After sonication was completed, the detected temperature rise decreased immediately versus dissipating over time. This would indicate the detected thermometry was due to a phase change of the tissue rather than temperature increase. During treatment, a high amount of lower broad band emissions at frequency < 1.2MHz were detected by a cavitation sensor. Post-treatment MR imaging showed that at power levels between 300 and 400 W, there were areas of hypointensity indicating the lesion. H&E staining confirmed the presence of the mechanical lesion where various anatomical targets were fractionated. Power levels were sufficient to rupture vessels and cause a focused area of haemorrhage at the treatment cluster. It showed that BH treatment dissolved the anterior ventricle wall with presence of elements of blood. H&E staining also showed that maximum lesion diameter was approximate 7 mm in coronal plane therefore the treatment borders matched the treatment plan. For the piglet treated with 1-ms long pulses, the post-treatment imaging change was not noticeable. However, at 375 and 400 W power, H&E slides showed two areas where there was a perforation of the anterior ventricle wall with a lesion size of up to 2 mm. It appeared that the shorter pulse duration generated smaller but more focused lesions.

Conclusions

This pilot study shows that the clinical Sonalleve system is capable of generating mechanical ablation of a brain tissue in an in vivo porcine model using boiling histotripsy pulsing scheme. The power threshold to initiate lesions in brain using 10 ms pulses (375 W) was found to be similar the power levels used in BH studies in ex vivo bovine liver and porcine kidney tissue at a similar depth in tissue (250 – 300 W) in a Sonalleve V2 system at the University of Washington. The treatments can be accelerated by using higher power outputs and shorter pulses. H&E histological evaluation showed that BH treatment caused rupture of vessels focally while also creating wider well defined areas of mechanical ablation with no damage to surrounding tissue. Additional work is underway to characterize the pressure levels generated by the Sonalleve to correlate the power and pressure for treatment.
Fig. 7 (abstract O15).

Top Row: Left to Right: Treatment area, Pre T2-w (100–500 W), Post T2-w (100–500 W), Pre- T2-w (325–425) and Post T2-w (325–425); Bottom Row: Histology Left to Right: 10 ms treatment (375W), 10 ms treatment - zoomed (375W), 1ms treatment (375W and 425W), 1 ms treatment – zoomed.

O16 Musculoskeletal clinical applications of intense therapy ultrasound (ITU): part 1. Clinical study for plantar fasciitis

Michael Slayton1, Richard C. Amodei1, Keegan Compton1, Ashley McNelly2, Daniel Latt2

1Guided Therapy Systems, LLC, Mesa, Arizona, USA; 2School of Medicine, University of Arizona, Tucson, Arizona, USA

Objectives

Chronic Plantar fasciitis (CPF) is a common cause of plantar heel pain that is a result of a degenerative process of the plantar fascia and its surrounding perifascial structures [1]. It is the most common cause of heel pain, affecting 10% of the U.S. population, and one of the most common foot and ankle problems.

More than twenty different treatments have been used for plantar fasciitis. Conservative treatments (rest, ice, stretching and NSAIDs) have been shown to effectively treat symptoms but 10% of patients fail conservative management and continue to have symptoms within 12 months and beyond. Surgery consisting of partial PF release is often considered with 50% of patients having residual symptoms, in addition to surgical risk exposure.

High frequency ITU, a novel potential approach to treating Plantar Fasciosis, was studied for the creation of small thermal injuries noninvasively inside symptomatic Plantar Fascia (PF). It has been shown to initiate a tissue repair cascade and promote collagen generation in musculoskeletal tissue [2, 3]. A double blinded, randomized, sham controlled clinical study for ITU treatment of chronic Plantar Fasciitis has been conducted by IRB approved clinical protocol to access clinical efficacy of the procedure.

Methods

Custom 3.2 MHz high intensity (10 kW/cm2) ultrasound therapy system was designed and fabricated (GTS, Mesa, AZ, USA). Field simulations, testing and Schlieren images verified intensity, high focal pressure (17.3 MPa) and focal distance of 13–15 mm.

Each treatment consisted of 250–320 100 ms pulses creating matrices of small ablative thermal lesions of 4–5 joules at pre-programmed pitch of 1.6 mm. Each patient underwent two treatment sessions in 2 weeks, each treatment time did not exceed 12 minutes. ITU placebo group consisted of the same treatment with energy set to 0.

Treatment effects were assessed with diagnostic imaging ultrasound at 12 MHz (Spark, Ardent Sound, Mesa, AZ, USA) by a certified sonographer. Ultrasound images were analysed to determine symptomatic hypoechoic lesion size with PF.

Patient reported outcomes consisted of PROMIS physical function computer adaptive test (PF-CAT), PROMIS global health, Foot Function Index pain subscale (FFIPS) [4, 5] and a non-validated heel pain specific questionnaire.

Clinical protocol included (35) patients diagnosed with chronic heel pain due to Plantar Fasciitis (more than 3 months) and failed conventional therapy treatments.

Patients were randomized to standard therapy (anti-inflammatory pills, stretching and gel heel cups) plus ITU (“Treatment” group, n=26) of standard therapy plus sham ITU (“Control” group, n=9).

Primary investigator, sonographer and study coordinator administering the study were blinded to group assignments. P-values were calculated via 2-tailed paired T-tests for both treatment and control groups.

Results

Patient-Reported Outcome Measures: Compared to the baseline assessment of Pain, the Treatment Group showed significantly improved pain scores compared to the Control (sham treatment) Group in follow-up visits including 12 weeks after the initial treatment.

Foot Function Index Pain Score: Compared to the baseline assessment, the treatment group pain scores also showed significant improvement compared to the sham group.

Diagnostic Ultrasound Imaging: During the 12 week follow-up period changes to the overall thickness of the PF were not statistically significant, while calculated volume size of hypoechoic lesions within the PF, just distal to the Calcaneus, showed significant change. For the experimental group (n=28) the average hypoechoic lesion volume reduction was followed and compared to the baseline measurements just before the first Treatment; 2 week follow-up and 2nd treatment date (−28%), 4 weeks (−50%), 6 weeks (−66%) and 12 weeks (−80%).

For the control group (n=10), the average hypoechoic lesion volume was followed and compared to the baseline just before the first Treatment; 2 weeks and 2nd sham treatment (+9%), at 4 weeks (+16%), 6 week (+29%) and 12 weeks (+31%). Unlike the experimental group, these lesions grew in size during the follow-up period.

P-values calculated for all outcome results discussed above for both treatment and control groups were below .01, showing the statistical significance of the results.

Conclusions
  1. (1)

    Results of the double blinded randomized, sham controlled study for the treatment of Plantar Fasciitis with ITU appeared to have statistically significant positive results within 12 weeks post-treatment in 80% of treated subjects.

     
  2. (2)

    Both quantitative measurements from diagnostic ultrasound imaging and applied standardized assessment protocols consisting of PROMIS PF-CAT, FFIPS along with Patient Reported Outcome Measures showed statistically significant coincidental improvements in treated subjects vs. control group.

     
  3. (3)

    Intense Therapeutic Ultrasound has shown potential for effective treatment of Chronic Plantar Fasciitis. Better designed studies with increased # of subjects will be considered to support ITU as an effective tool for the proposed clinical treatment.

     

References

[1] Neufeld S.K. and Cerrato R.; Plantar fasciitis: evaluation and treatment. J Am Acad Orthop Surg, June 2008, 16(6): p. 338–46.

[2] Slayton M. and Barton J.; Healing tissue response with ITU (Intense Therapy Ultrasound) in musculoskeletal tissue, feasibility study. Ultrasonics Symposium (IUS), 2014 IEEE International, Chicago, USA, p.1654–1657.

[3] Slayton M. H., Amodei R. C., McNelly A. and Latt D. L.; Intense Therapy Ultrasound (ITU) for the treatment of Chronic Plantar Fasciitis: Preliminary Results of Clinical Study. 37th International Conference of the IEEE Engineering in Medicine and Biology; Milan, Italy, August 2015.

[4] Rose, M., Bjorner, J.B., Becker J., Fries J.F. and Ware J.E.; Evaluation of a preliminary physical function item bank supported the expected advantages of the Patient-Reported Outcomes Measurement Information System (PROMIS). J. Clin. Epidemiology, January2008, 61(1): p. 17–33.

[5] Budiman-Mak E., Conrad K. J. and Roach K. E.; The Foot Function Index; a measure of foot pain and disability. J. Clin. Epidemiology, 1991, 44(6): p. 561–579.
Fig. 8 (abstract O16).

Treatment Group SROM 1

Fig. 9 (abstract O16).

Control Group SROM 1

Fig. 10 (abstract O16).

FFIP Score by Visit

Fig. 11 (abstract O16).

Average Lesion Volume change by Visit. SE applied.

O17 Musculoskeletal clinical applications of intense therapy ultrasound (ITU): part 2. Initial results of clinical study for lateral epicondylitis

Michael Slayton1, Richard C. Amodei1, Keegan Compton1, John Kearney2

1Guided Therapy Systems, Mesa, Arizona, USA; 2The CORE Institute, Phoenix, Arizona, USA

Objectives

Acute and Chronic pain of the Common Extensor Tendon (CET) region, lateral epicondylitis or tennis elbow is a common pathology of both athletes and non-athletes affecting up to 3% of the population at large [1], while the prevalence of chronic problems caused by overuse in tennis players can be as high as 40%. Elbow tendinopathy represents an important set of pathologies that account for lost recreation time, decreased quality of life, and work-related disability claims. Conservative treatment of the Epicondylitis or -osis is recommended as the initial strategy by most authors. This strategy includes identification and correction of possible etiological factors, and a symptom related approach. Generally, the initial treatment consists of a multifactorial approach that may include a combination of rest (complete or modified activity), medication (NSAIDs for Epicondylitis), stretching and strength training. More aggressive treatments for CET include: Cortisone injection, Plasma Rich Platelets (PRP), Tenotomy, ESWT.

High frequency ITU, a novel potential treatment for CET, was studied for the creation of small thermal injuries noninvasively inside symptomatic Common Extensor Tendon (CET). It has been shown to initiate a tissue repair cascade and promote collagen generation in musculoskeletal tissue [2–5]. A blinded, randomized, clinical study for ITU treatment of chronic Lateral Epicondylitis has been conducted by IRB approved clinical protocol to assess clinical efficacy of the procedure.

Methods

Custom 4.5 MHz high intensity (47.9 KW/cm2) ultrasound therapy system was designed and fabricated (GTS, Mesa, AZ, USA). Field simulations, testing and Schlieren images verified intensity, high focal pressure (37.9 MPa) and focal distance of 6 mm.

Each treatment consisted of 80 14 ms pulses creating matrices of small ablative thermal lesions of 1 joule at manually targeted area set by diagnostic ultrasound imaging. Each subject underwent two treatment sessions 4 weeks apart. Each treatment time did not exceed 10 minutes. Treatment effects were assessed with diagnostic imaging ultrasound at 20 MHz (Spark, Ardent Sound, Mesa, AZ, USA) by a certified sonographer. Ultrasound images were analysed to determine changes in the peri-tendon region, including hypoechoic areas, calcifications and dependent free fluid.

Subject reported outcomes consisted of PRTEE survey [6], physical examination, Universal Analog Visual Pain Score17 and a Patient Reported Satisfaction Survey [7].

Clinical protocol includes 25 subjects diagnosed with chronic Tennis Elbow, or Lateral Epicondylosis (more than 3 months) and failed conventional therapy treatments.

Subjects were subjected to standard therapy (stretching and strength exercises, hot and cold compresses and compression support) plus ITU.

Primary investigator, sonographer and study coordinator administering the study were blinded to group assignments. P-values were calculated via 2-tailed paired T-tests at each visit of the clinical study.

Results

The results presented below are initial findings for the first 12 subjects currently being followed through the study.

PRTEE: Patient Reported Tennis Elbow Evaluation Final Score is a weighted Pain Score based on 15 questions grouped into 3 categories: Overall Pain, Functional Disability and Usual Activities. Subjects respond to each question with a Pain Score of 0–10. Each category is then summed and weighted with a maximum score of 100 (Overall Pain 100%, Functional Disability 50% and Usual Activities (50%), n=12, Fig. 12.

Self-Reported Outcome Measures Surveys show a significant improvement and treatment satisfaction with Subjects reporting improvements in elbow pain 100%, improvement in Daily function 83% (vs. 17% no improvement) and treatment satisfaction 83% (vs. 17% not satisfied), n=6, Fig. 13.

Universal Analog Pain Scores also show progressive reduction (−3 on a 10 point scale) throughout the same period, n=12, Fig. 14.

Diagnostic Ultrasound Images: Diagnostic Ultrasound Images show a consistent increase of free fluid 2 weeks after the first treatment, with a progressive reduction in free fluid at 8 weeks in subjects with no to mild peri-tendon calcifications. These subjects correlated well with improving PRTEE survey scores. Subjects with little or no improvement in PRTEE scores consistently presented with moderate to severe peri-tendon calcifications.

P-values calculated for the above reported outcomes were not statistically significant for Visits 2 and 3 (P>0.05) while results for Visits 4 and 5 demonstrated P< 0.05.

Conclusions
  1. 1.

    Feasibility of Intense Therapeutic Ultrasound treatments of chronic pain in CET region has been established with the initial results (n=12) of the ongoing clinical study.

     
  2. 2.

    Significant reduction of pain scores per activities (PRTEE) and Self-Reported Outcome Measures (83% improvement) with average Universal Pain Scores reduction from 5.0 to 2.0 were statistically significant (p<0.05) at 8 and 12 weeks post treatment.

     

References

[1] Hong QN, Durand MJ, Loisel P. Treatment of lateral epicondylitis: where is the evidence? Joint Bone Spine 2004; 71(5):369–373.

[2] White, W. M., I. R. Makin, et al. (2007). Selective creation of thermal injury zones in the superficial musculoaponeurotic system using intense ultrasound therapy: a new target for noninvasive facial rejuvenation." Arch Facial Plast Surg 9(1): 22–29.

[3] Gliklich R, White WM, Barthe PG, Slayton MH, Makin IRS. Clinical pilot study of intense ultrasound (IUS) therapy to deep dermal facial skin and subcutaneous tissues. Arch Facial Plast Surg 2007; 9:88–95.

[4] Slayton M., Barton J, Feasibility of Modulating Healing Tissue Response by ITU (Intense Therapy Ultrasound) in Musculoskeletal Tissue ASLMS 2014 Annual Conference.

[5] Slayton M., Barton J., Healing tissue response with ITU (Intense Therapy Ultrasound) in musculoskeletal tissue, feasibility study, Ultrasonics Symposium (IUS), 2014 IEEE International, pp. 1654–1657. DOI 10.1109/ULTSYM.2014.010

[6] Rompe JD1, Overend TJ, MacDermid JC. Validation of the Patient-rated Tennis Elbow Evaluation Questionnaire.J Hand Ther. 2007 Jan-Mar; 2007 (1):3–10: quiz 11.

[7] Rose, M., et al., Evaluation of a preliminary physical function item bank supported the expected advantages of the Patient-Reported Outcomes Measurement Information System (PROMIS). J Clin Epidemiol, 2008. 61(1): p. 17–33.
Fig. 12 (abstract O17).

See text for description

Fig. 13 (abstract O17).

See text for description

Fig. 14 (abstract O17).

See text for description

O18 Clinical experience of intra-operative high intensity focused ultrasound in patients with colorectal liver metastases. Results of a phase ii study.

David Melodelima1,2, Aurelien Dupre2, Yao Chen2, David Perol2, Jeremy Vincenot1, Jean-Yves Chapelon1, Michel Rivoire2,1

1LabTAU - U1032, INSERM, Lyon, France; 2Centre Leon Berard, Lyon, France

Objectives

Managing colorectal liver metastases (CLM) is a major clinical challenge, and surgery remains the only potentially curative treatment. Five-year survival rates of up to 51% have recently been reported. However, only 10–20% of patients are eligible for surgery, which is often precluded by the number, size and/or location of metastases, or because the necessary resection will leave insufficient volume of functional liver. Radiofrequency ablation (RFA) is the main technology that has been used in association with surgery as a tool to expand the number of patients who may be candidates for liver directed therapy. However, RFA has several limitations. There is a risk of inadequate treatment due to the heat sink effect of blood flow, RFA does not allow reliable real-time monitoring, and it require intra-parenchymal introduction of a probe. Moreover, only small hepatic volumes can be targeted. These limitations could explain the high rates of local recurrence seen after RFA.

High intensity focused ultrasound (HIFU) has been proven effective in a wide range of clinical applications, especially prostate cancer. The ablation achieved by conventional HIFU is small and ellipsoidal. The dimensions vary according to transducer characteristics but are typically 1–3 mm (transverse) and 8–15 mm (along beam axis). In clinical practice, hundreds of superimposed ablations are required; and the procedure may take up to two hours. Even so, HIFU has several potential advantages in the treatment of liver tumours: there is no need to puncture the parenchyma, the extent of the thermal lesions achieved is not reduced by hepatic perfusion, and it is possible to monitor the effects of therapy in real time. However, extra-corporeal treatment of the liver is difficult because presence of the ribcage may stop propagation of ultrasound waves and respiratory motion may cause targeting problems. HIFU treatment of CLM needs to be improved, and reducing the duration of surgical intervention by increasing the size of ablated fields is particularly important. A HIFU device enabling destruction of larger liver volumes has been developed based on toroidal transducers.

Preliminary in vitro and preclinical work demonstrated the potential, feasibility and safety of such HIFU ablations. During laparotomy in a porcine model we demonstrated that this HIFU device achieves reproducible ablations with an average volume of 7 cm3 (with 20 mm diameter and 25 mm long axis) in 40 seconds. Such preclinical work has to be translated into clinical practice through controlled trials, and the aim of this study was to assess the feasibility and safety of HIFU ablation in patients undergoing hepatectomy for CLM, as well to collect efficacy and accuracy data. This study is registered with Clinical-Trials.gov (NCT01489787).

Methods

This study was a prospective, single-centre phase I/II study evaluating the feasibility, safety and accuracy of HIFU during surgery in patients with CLM. The protocol was reviewed and validated by a national ethics committee (CPP Sud-Est IV) according to French and European directives. Since this study was the first use in man of intra-operative hepatic HIFU, ablations were made only in areas of liver scheduled for resection. This allowed real-time evaluation of HIFU ablation while protecting participating patients from any adverse effects related to this new technique. The transducer has a toroidal shape 70 mm in diameter and is divided into 32 ultrasound emitters of 0.13 cm2 operating at 3 MHz. The radius of curvature is 70 mm. A 7.5 MHz ultrasound imaging probe was placed in the centre of the device and was used to guide the treatment. The imaging plane was aligned with the HIFU focal zone.

Six patients were included in the Phase I. Two single thermal ablations were created in each patient. Thirteen patients were included in Phase IIa and two HIFU ablations were to be placed precisely in a target previously identified in ultrasound images (step 1) and then at distance (step 2) from a target. Five patients were included in Phase IIb until now. HIFU ablations were created to ablate metastases (20 mm maximal diameter) with safety margins in all directions. The exposure time varied from 40 s to 370 s according to the diameter of the metastases to be treated.

Results

In agreement with preclinical studies, the demarcation between ablated and non-ablated tissue was clearly apparent in ultrasound images and histology. The dimensions measured on ultrasound imaging were correlated (r=0.88, p<0.0001) with dimensions measured during histological analysis. All HIFU ablations were obtained in 40 seconds. The average dimensions obtained from each HIFU ablation were a diameter of 21.0 ± 3.9 mm and a depth of 27.5 ± 6.0 mm. The phase IIa study showed both that the area of ablation could be precisely targeted on a previously implanted metallic mark and that ablations could be created deliberately to avoid such a mark. Ablations were achieved with a precision of 1–2 mm. In Phase IIb, one metastasis of 10 mm in diameter was ablated in 40 seconds with safety margins. Using electronic focusing metastases of 2 cm in diameter were ablated with safety margins (>3 mm in all directions) in 370 seconds. The dimensions of HIFU ablations created in 370 s were a diameter of 48 mm and a long axis of 51 mm.

Conclusions

This new HIFU device safely achieved large volume liver ablations in 40 s, with a precision of one to two millimetres under real-time monitoring. HIFU ablations of small metastases (<20 mm) and peri-lesional healthy liver were successfully created with planned safety margins of at least 3 mm in all directions.

O19 Chemotherapy in oral cancer

Wei Guo, Guoxin Ren, Guofeng Shen

Department of Oral Maxillofacial and Head Neck Oncology, Shanghai 9th People's Hospital, Shanghai, China

Objectives

To evaluate the efficacy and the main side-effects during the clinical trial of this new ultrasound hyperthermia system combined with chemotherapy in oral cancer, meanwhile, to observe the preliminary clinical response of this combined therapeutic modality.

Methods

Thirty four cases of oral squamous cell carcinoma entered this clinical trial, from which 23 had advanced oral carcinoma and were treated with new ultrasound hyperthermia system combined plus docetaxel–cisplatin–fluorouracil regimen (test group). Eleven patients only received chemotherapy with docetaxel–cisplatin–fluorouracil regimen (control group). The thermo-index were detected during the course of hyperthermia, the chief-complain of the patients were also recorded. The systemic physiological, biochemical and immunological index were tested before and after the treatment respectively. The therapeutic response was estimated 1 month after 2 cycles of the treatment.

Results

Twenty three cases of oral squamous cell carcinoma enrolled the clinical trial of local ultrasound hyperthermia combined with chemotherapy. 230 times of ultrasound hyperthermia in total were performed. The ultrasound hyperthermia system operated smoothly, no malfunction was found. The main thermo-index were: the maximum heating temperature was108 F, the average heating temperature was 106 F, the minimum heating temperature was 104 F, the fraction of heating time more than 108 F was 0.46, the average treatment time was 37.74±8.88min. PR+CR was 71% (test group). The main local side-effects were low-grade pain (6/23). The incidence of adverse effects was similar between both study groups, no bone marrow suppression (over III).

Conclusions

The system combined with docetaxel–cisplatin–fluorouracil regimen is effective and safe in the treatment of advanced oral cancer. The main side-effects of local ultrasound hyperthermia combined with chemotherapy are low-grade pain or tolerable pain. There is no immune function and obtains satisfying short-term response.

O20 Non-thermal, non-cavitational, 20kHz Ultrasound applicators in wound healing

Michael Neidrauer1, Leonid Zubkov1, Michael S. Weingarten2, David J. Margolis3, Peter A. Lewin1

1The School of Biomedical Engineering, Science and Health Systems, Drexel University, Philadelphia, Pennsylvania, USA; 2Department of Surgery, College of Medicine, Drexel University, Philadelphia, Pennsylvania, USA; 3Biostatistics and Epidemiology, University of Pennsylvania, Philadelphia, Pennsylvania, USA

Objectives

This talk examines the challenges associated with the design of clinically viable ultrasound applicators operating at the relatively low frequency (20 kHz) and intensity (<100 mW/cm2, spatial peak, temporal peak) levels, and tailored to treatment of chronic wounds, such as venous or diabetic ulcers. These challenges were associated with the architecture and weight, and principle and efficiency of operation, including electrical power consumption. The ultimate goal of this work was to test the efficacy of the applicators in human subjects.

Methods

A fully wearable Band-Aid™-like, dial-in delivery, battery-operated ultrasound applicator was designed. The applicator included light weight (<25g) piezoelectric flexural transducer and was powered by 10-12V fully rechargeable lithium-ion batteries (total weight <200g); it was able to operate for up to 4 hours between re-charging. To emphasize the uniqueness of the design, it might be useful to note that typically, the thickness of the capacitive piezoelectric element is inversely proportional to the frequency, therefore a 20 kHz element would need to be 10 cm thick. Such element would be bulky and require hundreds of volts (demanding a large power amplifier), in excitation signal thus eliminating any chance of being a portable design. To overcome this, a mechanical displacement amplifier, which translates 2 MHz ultrasound waves into 20 kHz output at the desired pressure amplitude (55 kPa; i.e. 100 mW/cm2) with only 12 volts excitation was chosen as a preferred solution.

The applicators were extensively tested to ensure that the ultrasound field energy was below the level needed to generate inertial cavitation and any temperature elevation that would exceed 1°C. Also, the uniformity of the acoustic field distribution was verified. The pilot study included 32 individuals between ages of 18 and 80 having venous (n=23) or diabetic (n= 9) wounds that remained open for a minimum of 8 weeks. In compliance with the IRB study protocol the subjects were randomly assigned to either treatment or control group, with an equal chance of being assigned to receive active ultrasound treatment or sham (current standard care). Treatment sessions lasted 15 minutes and were administered once a week for a period of 12 treatments, or until the wound’s closure. Clinical efficacy was evaluated by measuring the reduction in wound area over time. For both etiologies, i.e. both venous and diabetic wounds the rate of closure was statistically faster (p<.05) in the treated group compared to the control group.

Results

The study findings show that the ultrasound treated venous ulcer group had statistically improved (p<0.04) rate of wound size change (reduction of 14.3%/week) compared to the rate of wound size change for the control group (increase of 3.6%/week on average). Diabetic wound closure was achieved typically after 4 sessions for treated wounds, as opposed to 7 sessions for the control group. Time to heal was also statistically faster (p< .05) for treated wounds (~5 weeks) when compared to non-treated wounds (~12 weeks).

Conclusions

Overall, the results from this study support the notion that low frequency ultrasound treatment can successfully improve healing outcomes in chronic wounds with different morphology and etiology. The evaluated device used safe levels (<100mW/cm2 ISPTP) of ultrasound energy and featured unique portability, which opens possibility for personalized home treatment of chronic wounds in the future.

Acknowledgements

NIH Grant R01 EB9670, NSF 1064802, Wallace H. Coulter Foundation.

O21 Tests of thermal ablation with a 230 kHz transcranial MRI-guided focused ultrasound system in a large animal model

Nathan McDannold1, Jonathan Sutton1, Natalia Vykhodtseva1, Margaret Livingstone2

1Radiology, Brigham and Women's Hospital, Boston, Massachusetts, USA; 2Neurobiology, Harvard Medical School, Boston, Massachusetts, USA

Objectives

Thermal ablation with transcranial MRI-guided focused ultrasound (TcMRgFUS) is being tested clinically for as an alternative to surgery for functional neurosurgery and brain tumour resection. The current TcMRgFUS system, which operates at 650–670 kHz, is limited by skull heating to a small central region in the brain. Use of a lower acoustic frequency will reduce skull heating, but at the same time the focal heating will decrease and the risks of uncontrolled cavitation (the formation of microbubbles) increase. The purpose of this study was to evaluate the feasibility of thermal ablation in nonhuman primates using a system that operates at a lower acoustic frequency and to determine whether it can increase the “treatment envelope” for TcMRgFUS.

Methods

The experiments were approved by our institutional animal committee. Thermal ablation with the 230 kHz ExAblate Neuro system (InSighTec) was tested over five sessions in three rhesus macaques. In each session a target in the thalamus was sonicated transcranially at 40–50 s at acoustic power levels ranging from 90–560 W. The TcMRgFUS system software modulated the acoustic power in real time with a closed-loop controller that maintained a low-level of acoustic emissions, which are correlated with cavitation activity. MR temperature imaging (MRTI) was acquired at 3T (LX, GE) in a single plane using a 14 cm surface coil (TR/TE: 29/13 ms; flip angle: 30°). Measurements of the peak temperature rise at the focus and on the outer brain surface were compared for the different animals as a function of the applied acoustic energy. For the brain surface we measured the average temperature at the hottest 5% of the voxels and over a two voxel wide strip and we also normalized the measurement by the outer skull area. Temperature measurements were used to calculate the accumulated thermal dose, which was then compared to post-sonication T2-weighted, T2*-weighted, and contrast-enhanced T1-weighted MRI. The focal and skull-induced heating on the brain surface were compared to an earlier study performed in macaques with a 650 kHz version of this system.

Results

Focal heating sufficient to create an MRI-evident thermal lesion was achieved in 4/6 targets; the peak thermal dose exceeded 240 CEM43°C at these targets (Fig. 15). Heating at the focus was slightly higher than that measured on the brain surface. The focal heating increased linearly as a function of the applied energy at a rate of 3.2 ± 0.4°C per kJ (R2: 0.81). The surface area of the outer skull ranged from 47–55 cm2. For the hottest 5% of the voxels on the brain surface included in the MRTI imaging plane, the temperature rise increased linearly as a function of temperature at a rate of 126.6 ± 7.3°C per kJ/cm2. For a two voxel wide strip over the entire brain surface, this increase was 62.7 ± 7.5°C per kJ/cm2. The extent of MRI-evident changes (apparent oedema in T2-weighted MRI, BBB disruption post-contrast, no petechiae in T2*-weighted MRI) were consistent with 240 CEM43°C contours. One lesion imaged one week after FUS increased in size.

Conclusions

Analyses of the MRTI and post-sonication MRI suggest that the lesions were consistent with thermal mechanisms. The temperature rise increased linearly with the applied energy, and no evidence of cavitation-related petechiae were evident after sonication. The MRI-evident lesions were consistent with isodose contours drawn at 240 CEM43°C, a conservative threshold often used to guide thermal ablation. However, since it is known that thermal damage can take several hours to manifest in MRI and the lesion we imaged at one week increased in size, it is likely that the size of the lesion was underestimated by this dose value.

Similar tests in macaques with a version of this system operating at 670 kHz (Hynynen et al., Eur J Radiol 2006; 59: 149–56) measured skull-induced heating of 130°C per kJ/cm2 of outer skull surface, more than twice of that measured here (63°C per kJ/cm2). While no or minimal focal heating was observed at 670 kHz, with this 230 kHz system we were able to reach ablation-level thermal dose values. Thus, these preliminary results thus suggest that this low frequency system can expand the area of the brain that can be targeted for thermal ablation without overheating the skull. The closed-loop feedback system successfully maintained a low level of acoustic emission (and presumably microbubble activity) and immediately stopped the sonication when excessive levels were detected. However, additional work is needed to understand whether low-level cavitation activity played a role in the focal heating, to characterize the lesions in histology, and to examine whether safe cavitation levels can be maintained in tumours where the cavitation threshold may vary.
Fig. 15 (abstract O21).

MRTI (top) and post-FUS imaging (bottom) obtained in two sessions in Monkey 1. Thermal dose contours at 30 (orange) and 240 (red) CEM43C were calculated from the MRTI. Immediately after each session, a small lesion was observed in contrast-enhanced T1-weighted MRI (CE-T1WI). The dimensions of this area were consistent with the 240 CEM43 contours. The lesion produced in session 1 was largely non-enhancing in in CE-T1WI at week 2. It was visible in T2-weighted imaging (T2WI) and increased in size

O22 Growth slowdown in a brain metastasis model by antibody delivery using focused ultrasound-mediated blood–brain barrier disruption

Thiele Kobus1,2, Yong-Zhi Zhang2, Natalia Vykhodtseva2, Nathan McDannold2

1Radiology and Nuclear Medicine, Radboud University Medical Center, Nijmegen, Netherlands; 2Radiology, Brigham and Women's Hospital, Boston, Massachusetts, USA

Objectives

HER2-targeting antibodies prolong survival in patients with HER2-positive breast cancer metastases outside the brain. However, the response of brain metastases to these drugs is poor and it is hypothesized that the blood–brain barrier (BBB) limits drug delivery to the brain. We aim to improve delivery by temporary disruption of the BBB using focused ultrasound (FUS). Here we evaluate the treatment benefit of combining two antibody therapies that target the HER2-receptor with FUS-mediated BBB disruption in a breast cancer brain metastasis model.

Methods

MDA-MB-361 HER2-positive human cancer cells were injected in the right brain hemisphere of nude rats. The animals were divided in three treatment groups of 10 animals each: the control-group received no treatment; the antibody-only group received trastuzumab and pertuzumab (antibodies that target the HER2-receptor); and FUS+antibody-group received trastuzumab and pertuzumab in combination with FUS-mediated BBB disruption. The six weekly treatments started five weeks after tumour implantation, when the tumour diameter was around 2 mm. The FUS treatments took place in a 7T MR-scanner using a single-element, spherically-focused 690 kHz-transducer. Trastuzumab and pertuzumab were injected before the first sonication. At the start of each sonication (duration 60s, 10-ms bursts, burst repetition frequency 1 Hz), the ultrasound contrast agent Optison (100 μl/kg) was injected. The complete tumour was treated in 4 to 14 sonications that were separated 1 to 1.5 mm. Peak negative pressure amplitudes in water between 0.46 and 0.62 MPa were used.

Before and after the sonications, MR imaging was performed consisting of T2-weighted (T2w), T1w and T2*w imaging to locate the tumour, confirm BBB disruption and study the presence of hemorrhages. In two animals tumour leakiness was studied by comparing T1w imaging before and after gadolinium injection before the tumours were sonicated. The difference in signal intensity change in pre- and post-contrast T1w images was determined between the tumour and contralateral brain region (= SI%). In all FUS-treated animals BBB disruption was confirmed with contrast-enhanced T1w imaging and quantified using the same method as for the tumour leakiness. Pre- and post-sonication T2*w images were inspected for hypo-intense regions, which can indicate extravasated erythrocytes.

Every other week, high-resolution T2w imaging was performed to determine tumour volume. The growth rate (r) was determined by fitting the tumour volumes to the following formula: volume(t)=a*exp(r*t), in which t is the time in days. The growth rate of each tumour was determined for the treatment period (week 5 to 11) and the follow-up period (week 11 till sacrifice). An animal was classified as ‘responder’ if the growth rate was lower than the mean growth rate of the control animals minus two standard deviations.

The animals were euthanized if the tumour size exceeded 13 mm in diameter or if the condition of the animal was poor. From nine animals, histology was obtained (hematoxylin and eosin (H&E)) and the brains of five animals were stained for HER2.

Results

BBB disruption was successful in all sessions with an average SI% of 21.2% (range 4.5 – 77.6%). The mean SI% of two tumours before BBB disruption during the six treatment weeks were 0.4% and 0.6%, indicating that the tumours were not leaky before disruption (Fig. 16). In 33% (20/60) of the FUS-sessions, regions were present that were clearly more hypo-intense on post- than on pre-sonication T2*w images, suggesting hemorrhages. In the remaining 67% of the sessions, no or a small difference in hypo-intensity was observed.

In the FUS+antibody-group, 4/10 animals were classified as responders during the treatment period (week 5 to week 11) with an average growth rate of 0.010±0.007, compared to 0.043±0.013 for the non-responders. There was no difference in the average SI% of the responding rats (21.8%±16.7) and the non-responding rats (20.7%±9.7). None of the control or antibody-only animals were classified as responder. When the FUS+antibody-animals are grouped, no significant differences in mean growth rates were observed between the control, antibody-only and FUS+antibody animals for the treatment period, nor for the follow-up period.

High-resolution T2w imaging showed that the tumour was homogenous in almost all animals till week 13–15, when cystic and necrotic areas started to develop. The tumours showed also a heterogeneous appearance on H&E stained sections and the complete tumour was expressing the HER2-receptor in the examined animals.

Conclusions

In this study, we demonstrate that FUS-mediated BBB disruption in combination with antibody therapy can slow down the growth of brain metastasis from breast cancer. As the tumours were not leaky before BBB disruption and no difference in growth rates was observed in the antibody-only group, the disruption of the BBB is necessary for drug delivery to these brain metastasis. Interestingly, only part of the rats responded to the treatment, the other animals had the same growth rate as the control-group. This is in line with a previous study (Park et al. 2012, J. Control. Release), where antibody therapy was combined with FUS in a different brain metastasis model and in part of the animals a strong response was observed, while the other animals did not respond. We did not observe a difference in tumour volume at the start of the treatment, in HER2 expression on histopathology, or in contrast-enhancement on MR images between the responders and non-responders to explain this. Better understanding of why certain animals respond is needed and will help in translating this technique to the clinic.
Fig. 16 (abstract O22).

a T1weighted image before contrast administration. The red arrow indicates the tumour. b No difference in enhancement of the tumour is observed after contrast administration (SI=0.4%). c After focused ultrasound-mediated blood–brain barrier disruption, the tumour enhances after contrast administration (SI=30.1%)

O23 Long term follow up of 6 essential tremor patients treated with MR-guided focused ultrasound thalamotomy

Michael Schwartz1,2, Yuexi Huang2, Nir Lipsman1, Jennifer Jain2, Martin Chapman2, Tejas Sankar3, Andres Lozano1, Kullervo Hynynen2

1Surgery (Neurosurgery), University of Toronto, Toronto, Ontario, Canada; 2Sunnybrook Health Sciences Centre, Toronto, Ontario, Canada; 3Surgery (Neurosurgery), University of Alberta, Edmonton, Alberta, Canada

Objectives

To determine the factors influencing outcome after MR-guided focused ultrasound (MRgFUS) thalamotomy for essential tremor.

Methods

Between May 2012 and February 2013, 6 patients were treated with MRgFUS thalamotomy. The first 4 patients have been reported (Lipsman et al., Lancet Neurol. 2013 May; 12(5): 462–8). Prospective recording of data from preoperative screening to 3 month follow up according to an experimental protocol using the clinical rating scale for tremor (CRST), and then a 2 year follow up examination and assessment were done. Maximum temperature at the focus of sonication was determined and post-treatment day 1 MRI scans were reviewed for lesion size and location. A point derived by our current method of determining the probable location of the ventral intermediate nucleus (VIM) of the thalamus was taken as the starting point for each treated case and a deviation vector from that point to the centre of each lesion produced was plotted.

Results

For the 6 patients, the mean CRST A scale (rest, posture and action) for the treated hand and arm prior to treatment was 7.2. At 1 week post-treatment, the mean CRST A was 0.67. At 3 months, the mean CRST A was 0.83 and at 2 years approximately 3.33. (At 2 years, the CRST A for 2 patients was estimated from a narrative account). Mean lesion size, excluding one very large lesion with a volume 2.7 times the mean volume of the other 5 patients, was 90.2 mm3. Only 2 patients had no decline in thalamotomy effect. Of these, one had a relatively large lesion of 107.5 mm3, which was located at the predicted location of the VIM nucleus. The other patient’s lesion was located 2.4 mm lateral to the predicted VIM location. He had prominent ataxia following his treatment. This subsided by 3 months. The patient with the large lesion, treated for left hand tremor, had persistent tingling of the left side of his mouth, and his left thumb and 3 fingers but excellent early relief of tremor. Although there was some persisting reduction of tremor at 2 years, he could no longer write, nor drink without spilling. The centre of his lesion was 1.3 mm lateral and 0.3 mm posterior to the expected VIM location. In its superior-inferior dimension his lesion measured 9.4 mm. The maximum temperature achieved at treatment was 63°C, higher than the mean of 58°C. This decline may, in part, be due to the progress of his condition, with increasing tremor on his untreated side and the onset of tremor in both legs. Two of the patients with a decline in function had relatively small lesions.

Conclusions

Not all patients have lasting benefit from MRgFUS thalamotomy. The effect of relatively large lesions may be more durable, but lesioning temperatures of greater than 60°C should likely be avoided.

O24 Skull bone marrow injury caused by MR-guided focused ultrasound for cerebral functional procedures

Michael Schwartz1,2, Robert Yeung1, Yuexi Huang1, Nir Lipsman2, Jennifer Jain1, Martin Chapman1, Andres Lozano2, Kullervo Hynynen1

1Sunnybrook Health Sciences Centre, Toronto, Ontario, Canada; 2Surgery (Neurosurgery), University of Toronto, Toronto, Ontario, Canada

Objectives

To determine the factors that lead to bone marrow injury in the skulls of patients undergoing MRgFUS cerebral ablations for the treatment of movement disorders such as ventral intermediate (VIM) nucleus thalamotomy for tremor and pallidotomy for L-dopa induced dyskinesia (LID), and to follow the course of healing.

Methods

All patients undergoing functional cerebral ablations produced for the mitigation of movement disorders are followed with serial MRI scans according to study protocol to assess the evolution of the cerebral lesions. In one patient subjected to very high power sonication in the attempt to produce a pallidotomy for LID, skull lesions were noticed on a follow up MRI scan, produced 4 months after his MRgFUS procedure. A review of other treated patients is currently underway.

Results

Multiple ovoid lesions throughout the calvarium, new since the immediate post-treatment MRI scan done January 30, 2015, were seen on the MRI scan done May 12, 2015 (Fig. 17). Their appearance resembles that of bone infarcts. The MRI scan was repeated on October 2, 2015. Many of the ovoid lesions were still visible. This patient underwent sonication increasing to a maximum power of 1100 W for 31 seconds. During the procedure, the scalp and skull were constantly cooled with flowing degassed water at 13 °C. Despite this sonication, the target locus in the globus pallidus reached only 48 °C.

Conclusions

High power and duration sonication for functional cerebral lesions may cause injury to skull bone marrow. A review of all patients treated with MRgFUS for movement disorders is currently underway to determine whether there have been other cases, and to determine the threshold for bone marrow injury.
Fig. 17 (abstract O24).

See text for description

O25 In vitro study using MR-guided focused pulsed ultrasound for destroying clots using thrombolytic drugs

Christakis Damianou1, Nikolaos Papadopoulos2

1Cyprus University of Technology, Limassol, Cyprus; 2City University, London, United Kingdom

Objectives

In this paper an extensive study of using MR-guided focused pulsed ultrasound system is presented for the treatment of stroke using thrombolytic drugs in a model in vitro.

Methods

A single element spherically focused transducer of 4 cm diameter; focusing at 10 cm and operating at 1 MHz was used. The transducer was mounted in an MR compatible robotic system of 3 axes. The artery was modelled using a silicone tube. Tissue was modelled using agar-evaporation-silica gel. Coagulated blood was used to model thrombus. A thermocouple was placed in the thrombus in order to measure the thrombus temperature.

Results

The effect of power, presence of bubbles, temperature, presence of agar-evaporation milk-siligal gel, time of sonication, pulse repetition frequency, presence of standing waves, flow velocity were investigated. The goal was to maintain a temperature increase of less than 1°C during the application of pulsed ultrasound (called safe temperature). With the application of ultrasound alone or thrombolytic drug alone there was no notable destruction of the thrombus.

Conclusions

With the combination of ultrasound and thrombolytic drugs sufficient destruction occurred after 30 min, but the rate of destruction of thrombus (mL/min) is considered low. Thus, the clinical use of focused ultrasound for sonothrombolysis despite the full parametric study that we performed is considered pessimistic.

O26 Low frequency in-vivo cavitation mapping

Alexander Volovick1, Javier Grinfeld1, Yoav Levy1, Omer Brokman1, Eyal Zadicario1, Ori Brenner2, David Castel3

1INSIGHTEC, Tirat Carmel, Israel, 2Weizmann Institute, Rehovot, Israel, 3Sheba Medical Center, Ramat Gan, Israel

Objectives

Current trans-cranial Magnetic Resonance guided Focused Ultrasound Surgery (tcMRgFUS) treatments are limited to the targets that are located at the centre of the skull. Superficial and peripheral targets are prone to lower acoustic intensity reaching the focal point due to the geometry of the skull and the ultrasound transversal through it. A possible solution for this problem is reduction of the transmitting frequency. Reducing the frequency is a trade-off between higher skull penetration and lower tissue absorbing. In addition, lower frequency has a lower threshold to induce cavitation. Previously it was reported ([1]) that the transient cavitation threshold for the muscle tissue acts linearly with frequency as where Pth is the cavitation pressure threshold and f is the transmitting frequency. However this threshold was not reported to cause any unintentional pathological damage. In the presented work, different levels of acoustic intensities resulting in various occurrences of cavitation were applied to an in-vivo pig brain model, reaching various temperature levels and causing various pathological effects.

Methods

36 subjects underwent craniotomy and were sonicated using various ultrasonic parameters (ExAblate Neuro low frequency system, 220 Khz, by INSIGHTEC, Ltd). The temperature rise was measured using MR thermometry (1.5T GE HDx MRI by GE Healthcare, with an integrated Head Coil, by InSighTec) and cavitation signal was measured and recorded using integrated hydrophones. 20 subjects underwent 2 weeks follow up with post procedural MR imaging one, five, seven and fourteen days after the procedure. The brains were harvested, fixed in formalin and sliced to 3 mm slices. Macro-pathological slices for randomly selected subjects were also performed.

Results

Figure 18 represents the graph of temperature rise as a function of applied energy. Figure 19 represents the graph of cavitation activity as a function of maximal acoustic power. Figure 20 shows a follow up imaging for 2 lesions performed on a single subject, Fig. 21 is a macro pathology slide for the subject presented in Fig. 20. In general the temperature rise grew linearly with the applied acoustic energy; the cavitation activity was linearly dependent to the maximum applied acoustic power. For sonications that were in the central cavitation activity area as presented on Fig. 19 the lesions were well defined and increased in size till the 5th day follow up and then reduced in size becoming scarf tissue as presented on Figs. 20 and 21. Macro pathology revealed that tissue rapture is seen on the micro level; however it is well defined and limited within the lesion area.

Conclusions

Cavitation threshold levels that were observed in brain agree with the levels reported in literature. Additional levels of cavitation were observed and associated with effects on tissue as seen in MRI and histology. The data collected suggests cavitation levels that can be applied while keeping lesioning effect to the confined area and avoiding haemorrhages in tissue. Integrating a real time control over the level of cavitation and keeping the level below the safety threshold results in safe and effective tissue ablation in brain.
Fig. 18 (abstract O26).

Maximal average temperature as a function of effective energy. The graph is based on the data of 502 sonications

Fig. 19 (abstract O26).

Maximal cavitation activity as a function of effective power. The graph is based on the data from 253 sonications

Fig. 20 (abstract O26).

An example for two weeks follow up imaging for T2, T1 and T2 flair images presented. Two lesions are visible on the 1 and 5 days follow up images, single lesion is seen for the 7 days follow up and no lesion is visible at the 14 days follow up imaging. The lesions locations are marked by red circles at the 14 days follow up images in order not to hide the lesion and edemic tissue for the earlier follow up images. The lesion on the left was produced by the sonication reaching 58°C, whereas the lesion on the right reached temperature of 56°C

Fig. 21 (abstract O26).

Macro pathology of the slides presented on Fig. 20. Only one lesion was detected

O27 Real-time, transcranial passive cavitation mapping for monitoring of the focused ultrasound-induced blood–brain barrier opening in primates

Shih-Ying Wu1, Julien Grondin1, Wenlan Zheng1, Marc Heidmann1, Maria Eleni Karakatsani1, Carlos J. Sierra Sánchez1, Vincent Ferrera3, Elisa E. Konofagou1, 2

1Biomedical Engineering, Columbia University, New York, New York, USA; 2Radiology, Columbia University, New York, New York, USA; 3Neuroscience, Columbia University, New York, New York, USA

Objectives

Real-time cavitation monitoring during the focused ultrasound (FUS) induced blood–brain barrier (BBB) opening is crucial in assessing and controlling the BBB opening outcomes and safety. Currently, passive cavitation detection using a single-element detector for quantification of the stable and inertial cavitation doses has shown good correlation with the opening volume and the molecular delivery efficiency in nonhuman primates (NHPs). However, an off-line magnetic resonance imaging (MRI) is required for confirming the targeting or opening area in the brain after the FUS procedure. It is therefore essential to develop transcranial cavitation mapping in providing the spatial distribution of cavitation intensity in real time in order to precisely assess and control the BBB opening outcomes with accurate targeting during the procedure. The aim of this study is to develop real-time cavitation mapping using time-exposure acoustics and passive sparse matrix beamforming with the performance evaluated in both the in vitro primate skull (NHP and human) experiments and the in vivo NHP experiments during BBB opening.

Methods

Similar to array-based passive cavitation detectors, a linear probe (L7-4, Philips) with a programmable data acquisition system (Vantage, Verasonics) was used to acquire cavitation signal passively during the sonication (frequency: 0.5 MHz, pulse length: 10 ms, PRF: 5 Hz, pressure: 150–600 kPa) with in-house made microbubbles (lipid-shelled and monodisperse with a diameter of 4–5 μm) and a single-element FUS transducer with a coaxially aligned flat-band hydrophone as a single-element passive cavitation detector. Time-exposure acoustics for an integration of a series of passive cavitation images over the exposure time reconstructed by dynamic receive beamforming using sparse matrix calculation in graphic processing unit (GPU; Tesla K40, NVIDIA) were developed for the reconstruction of passive cavitation maps in real time. For the in vitro experimental setup, a phantom with a channel of 4 mm in diameter and an infusion pump was used to mimic the vessel with microbubble circulation (a concentration of 2×105 bubbles/mL with a flow rate of 1 mL/min), and the FUS transducer and the linear array were both focused at the channel orthogonally. B-mode imaging with the linear array was performed before acquiring the passive cavitation signal in order to confirm the alignment of both the linear array and the FUS transducer to the channel. Three sets of passive cavitation maps were acquired using this in vitro setup: 1) without the skull, 2) with the NHP skull and 3) with the human skull placed between the phantom and the linear probe. The effects of acoustic pressure, exposure time and aperture size to the intensity and focal size of the cavitation maps were all evaluated, as well as the computational cost in GPU and the sensitivity through the skull. Furthermore, the in vivo cavitation maps were acquired during the sonication for BBB opening in NHPs (duration: 2 min) and compared with the BBB opening outcomes in the MRI.

Results

The results of the in vitro experiments showed that the cavitation location corresponded to the microbubble disruption in B-mode images, and both the intensity and the focal area of the cavitation maps increased with pressure. Increasing the exposure time eliminated the interference outside of the focus and enhanced the focalization by minimizing the focal area, while both the intensity and the focal area reached a plateau at the exposure time of 62.5 μs. The focal size especially in the axial direction increased with decreasing aperture size, suggesting an improved focalization by using a larger aperture size of the array-based PCD. The computational time for the exposure time of 62.5 μs was 9.5 s, which can be decreased to 0.27 s in achieving real-time monitoring by decreasing the exposure time to 1.44 μs. By placing the skull, it was found that the cavitation signals were still detectable through the NHP and human skulls at 300 kPa and 600 kPa with the mapping system, respectively. For the in vivo experiments, the BBB opening in NHP were successfully monitored with passive cavitation mapping targeting at the caudate and the hippocampus, the deep subcortical structure in the brain.

Conclusions

A real-time cavitation mapping technique using time-exposure acoustics and passive sparse matrix beamforming has been developed with the performance and the sensitivity through the primate skull evaluated, and was used for monitoring of the BBB opening in NHPs. This novel transcranial monitoring technique providing both the spatial and intensity information of cavitation in real time during the FUS procedure is promising in assessing and controlling the targeting, treatment efficacy, and safety precisely.

O28 Blood brain barrier opening using focused ultrasound for the reduction of amyloid beta plaques in synergy with antibodies in a rabbit model fed with high cholesterol diet

Christakis Damianou, Marinos Yiannakou

Cyprus University of Technology, Limassol, Cyprus

Objectives

An animal model that creates Amyloid beta (Ab) plaques in the brain was implemented by delivering high cholesterol diet in rabbits for 3 months. The goal was to reduce the plaques using focused ultrasound (FUS) in combination with external antibodies.

Methods

A single spherically focused transducer was used which operates at 1 MHz, has focal length of 10 cm and diameter of 3 cm. The rabbit was placed in a custom made MRI compatible positioning device. Theoflavin staining was used in order to measure the plaque load at the end of each experiment.

Results

Using pulse FUS the blood brain barrier (BBB) was opened repeatedly up to 5 times at three day intervals. The opening of the BBB disruption was imaged using contrast-enhanced T1-weighted fast spin echo. By increasing the number of sessions, the number of plaques decreases (both for internal and external antibodies). With the use of FUS only (internal antibodies) the drop of average number of plaques/mm2 was reduced by 20% (in 5 sessions). The effect of external antibodies was more drastic. With 5 BBB sessions the average number of plaques/mm2 was reduced by 60%.

Conclusions

This study demonstrated that by opening the BBB, it will be possible to deliver internal and external antibodies to the brain, which eliminates Alzheimer disease (AD) plaques. More important by opening the BBB frequently (up to 5 times in this study) the reduction in the number of plaques is enhanced. Therefore FUS has the potentials to be used non-invasively for the treatment of AD.

O29 Correlation between down-regulation of p-glycoprotein and blood–brain barrier disruption in rat brain by mri-guided focused ultrasound and microbubbles

HongSeok Cho1, Hwayoun Lee1, Mun Han2, Jong-Ryul Choi1, Taekwan Lee1, Sanghyun Ahn1, Yongmin Chang2, Juyoung Park1

1Daegu-Gyeongbuk Medical Innovation Foundation, Daegu, Republic of Korea; 2Kyungpook National University, Daegu, Korea (the Republic of )

Objectives

Blood–brain barrier (BBB) is composed of both physical barrier with tight junctions and functional barrier with active efflux transporters. Mechanism of the functional barrier is mediated by P-glycoprotein (P-gp) and breast cancer resistance protein (BCRP) in brain endothelial cells. Past studies have shown that focused ultrasound (FUS) combined with microbubbles can disrupt the physical barrier of the BBB by exerting mechanical stress on the tight junctions; however, no study was performed to investigate the impact of FUS and microbubbles on the functional barrier of the BBB. Therefore, this study investigated the impact of BBB disruption induced by FUS and microbubbles on the expression of P-gp. We also investigated correlation between magnitude of the BBB opening and the down-regulation of P-gp.

Methods

A single target region was sonicated (0.5~0.7 MPa) transracially in one hemisphere in 3 rats using a 1 MHz FUS transducer; the other hemisphere served as a control. For the BBB disruption, 10 ms bursts were applied at 1Hz pulse repetition frequency (PRF) for 120 s and combined with IV injection of a microbubble ultrasound contrast agent (Definity 0.1 ml/kg). An MR contrast agent (Magnevist 0.4 mM/kg) and Evans Blue (0.15 ml/kg) were injected immediately after the sonication to indicate area of the BBB disruption in MR image and fluorescence spectroscopy, respectively. In order to measure the P-gp expression using a confocal fluorescence microscopy, the brains were fixed after perfusion and then stained immunohistochemically with a monoclonal antibody (C219) which reacts with a P-gp epitope.

Results

A T1 contrast enhanced MR image and Evans Blue fluorescent intensity at the sonicated regions indicated localized BBB disruption (Fig. 22). Both the MR contrast intensity and the Evans Blue fluorescent intensity were significantly increased in the targeted regions compared to the control regions (p<0.001). The fluorescence intensity of the P-gp expression at the confirmed locations of the BBB disruption was reduced by an average of 63.2±18.4% compared to the control area in all three rats. From the three sonicated regions, a total of 31 locations were selected and the P-gp fluorescence intensities were measured to observe the correlation between the degree of the BBB opening and the P-gp expression. Both the Evans Blue intensity and the MR contrast intensity were significantly correlated with the P-gp expression intensity (r=−0.72, p<0.001; r=−0.62, p<0.001, respectively). Histologic analysis on the sonicated region of the brain tissue revealed no apparent damage in the endothelial cells, and no significant amount of extravasated red blood cells was observed.

Conclusions

This study demonstrates that the BBB disruption induced by FUS and microbubbles reduces the expression of P-gp, and the level of the down-regulation of P-gp is significantly correlated with the magnitude of the BBB opening. These results suggest FUS + microbubble as a promising mean for the brain drug delivery through the BBB by overcoming both the physical and the functional barrier of the BBB.
Fig. 22 (abstract O29).

a A T1 contrast MR Image of a rat brain at sonification region (the opposite hemisphere of the brain served as control). b The fluorescence intensity of the P-gp expression at the soicated locations were reduced by an average of 63.2±18.4% compared to the control locations

O30 MR-guided focused ultrasound blood–brain barrier disruption through an intact human skull in a rat model using a clinical body system

Nicholas Ellens1, Ari Partanen1,2, Keyvan Farahani1,3, Raag Airan1

1Radiology, Johns Hopkins University, Baltimore, Maryland, USA; 2Philips, Andover, Massachusetts, USA; 3National Cancer Institute, Bethesda, Maryland, USA

Objectives

MR-guided focused ultrasound (MRgFUS) can be used for a large range of non-invasive therapies using mechanical and thermal mechanisms. For instance, it has been demonstrated that a combination of low intensity focused ultrasound and synthetic microbubble scan be used to safely, locally, and transiently disrupt the blood–brain barrier (BBB)[1].

Typically, transcranial MRgFUS thermal ablation requires large aperture phased arrays with many elements operating at lower frequencies and higher powers than body systems due to high transmission losses through the skull. However, in low pressure applications at low duty cycles, such as BBB disruption or neuromodulation, it may be feasible to use 'body' systems that have smaller aperture transducers with fewer elements [2]. The objectives of this study were to use a clinical body MRgFUS system to:
  1. 1)

    Quantify the ultrasound transmission and focal distortion through a human cranium at different orientations and in different focus locations inside the skull to assess the possible treatment envelope with such a device.

     
  2. 2)

    Demonstrate the ability to disrupt the blood–brain barrier (BBB) in rat neurovasculature through a human cranium.

     

Methods

A clean, degassed human cranium (gift of Dr. Quiñones-Hinojosa, JHU Neurosurgery) was mounted in degassed water above a 256-element phased-array transducer (14 cm focal length) of a clinical body MRgFUS system (Sonalleve V2, Philips, Vantaa, Finland). Hydrophone measurements were made with a lipstick hydrophone (Onda Corp., Sunnyvale, USA) mounted to a 3D stage, both in water and through the human skull in 11 locations (three orientations, 3–4 depths at each, ranging between the centre of the cranium and 2 cm from the skull surface). 40-cycle sonications at 1 MHz were applied at different acoustic powers ranging from 5 to 20 W (water) and 20 to 500 W (trans-skull). In each location, the full-width at half-maximum (FWHM) of the ultrasound focal point (in three dimensions) and its peak negative pressure were measured. This setup and the locations sampled are shown in Fig. 23.

Results

Hydrophone measurements demonstrated a great deal of variation in ultrasound transmission with changing transducer/skull incidence angle and skull thickness. The pressure attenuation ranged from −5.8 dB to −9.3 dB (mean +/−standard deviation, −8 dB +/− 1 dB), and the FWHM varied between 1.83 mm and 3.79 mm (2.2 mm +/− 0.3 mm in the anterior/posterior direction, 2.7 +/− 0.7 mm in the left/right direction). The insertion losses as a function of depth at different orientations are shown in Fig. 24. The dorsal midline skull (with the ultrasound beam passing through the sagittal and coronal sutures) was thicker than either of the angled approaches (passing through the parietal and temporal bones) and experienced a greater insertion loss for most sonication target depths. For both angled approaches, the insertion loss decreased as the focus was moved away from the skull surface. For the centered approach, the insertion loss increased slightly as the focus moved closer to the skull surface. It appears that the intervening skull thickness dominates the attenuation, though the angle of incidence seen by transducer elements also affects transmission.

Gadolinium enhancement in the brain on post-sonication T1-weighted MRI indicated successful BBB disruption. Disruption accuracy is shown in Figs. 25 and 26. For sonications of sufficient pressure to produce BBB disruption, the region of gadolinium enhancement did not appear to be shifted at all from the desired target. For higher estimated in situ pressures (>0.55 MPa), the disruption region was larger and less uniform, as evident in the 0.58 MPa point in Fig. 25. Even without element-by-element refocusing, the focusing quality and targeting precision appeared to be adequate. Post-experiment, there were no signs of gross damage to the animals to suggest significant off-target sonication.

Conclusions

Transcranial FUS focal point distortion was minimal despite the lack of element-by-element transducer refocusing. The clinical MRgFUS body transducer and driving electronics had sufficient power and aperture to generate the in situ pressures for BBB disruption through a human skull, using a variety of clinically practical approaches and patient orientations.

Though transcranial thermal ablation typically requires high acoustic powers, a lower operating frequency, and large aperture arrays with high number of elements, this study demonstrates that a clinical body MRgFUS system with a smaller transducer may be a safe and feasible alternative for non-invasive BBB disruption and other low-pressure therapeutic ultrasound transcranial applications, potentially offering a wide treatment envelope.

References

[1] Hynynen, K., et al. (2001). Radiology, 220(3), pp. 640–646.

[2] Airan, R. D., et al. (2015). 23rd ISMRM.
Fig. 23 (abstract O30).

a and b: Coronal and sagittal representations of the locations examined. c: Picture of the hydrophone and skull arrangement for one of the 'Left 45°' orientations. The transducer is below the table

Fig. 24 (abstract O30).

Insertion losses for different sonication orientations and depths

Fig. 25 (abstract O30).

Baseline T1-weighted coronal image, left, and post-contrast post-BBB disruption image, right, showing two locations of BBB disruption with the estimated in situ pressures labelled

Fig. 26 (abstract O30).

Sagittal images showing columns of BBB disruption produced by the labelled estimated in situ pressures. Ultrasound propagation is from the right to the left of these images.

O31 Interim results from a phase 1 clinical trial to disrupt the blood–brain barrier by pulsed ultrasound

Alexandre Carpentier3,4, Michael Canney1, Alexandre Vignot1, Cyril Lafon2, Jean-Yves Chapelon2, Jean-yves Delattre4,5, Ahmed Idbaih5

1CarThera, Lyon, France; 2INSERM, U1032, LabTau, Lyon, France; 3Assistance Publique Hopitaux de Paris, Hopital de la Pitie Salpetriere, Department of Neurosurgery, Paris, Paris, France; 4Universite Paris, UPMC, Paris, France; 5Assistance Publique Hopitaux de Paris, Hopital de la Pitie Salpetriere, Department of Neuro-Oncology, Paris, France

Objectives

Pulsed ultrasound, coupled with peripheral injection of microbubbles, has been shown in pre-clinical studies to be an effective method for enhancing the delivery of chemotherapy to the brain. In this work, an intra-skull implantable ultrasound device, SonoCloud®, was developed for temporarily disrupting the BBB. The device was implanted in patients with recurrent glioblastoma (GBM) undergoing systemic carboplatin chemotherapy in a Phase 1 clinical trial and the safety of repeated BBB disruption was assessed.

Methods

A Phase 1 clinical trial began in July 2014 at a single centre at the Hospital Pitie Salpetriere in Paris, France. Patients with recurrent GBM with an enhancing volume of less than 35 mm in diameter were included in the trial. Participants were implanted with a 11.5-mm diameter biocompatible 1 MHz ultrasound transducer, which was fixed into the skull bone thickness. The device was either implanted during a regular surgical resection of the enhancing region or during a unique ambulatory procedure under local anaesthesia. Once a month, the device was connected to an external radiofrequency generator using a transdermal needle, and patients received a two minute pulsed ultrasound exposure in combination with systemic administration of an ultrasound contrast agent (7 min mean total duration procedure). BBB disruption was assessed immediately after sonications using dynamic T1-weighted imaging with a gadolinium based MR contrast agent. Systemic intravenous injection of carboplatin chemotherapy was delivered immediately following MR imaging. Patients followed a progression of ultrasound dose in which the pressure was increased from 0.5 to 0.8 MPa throughout the course of the study.

Results

As of July 2015, eleven patients had been included in the study with a total of 25 BBB disruption procedures performed. No adverse effects were observed in patients treated. BBB opening was clearly observed in 12/25 treatments on contrast-enhanced T1w imaging. The procedure was safely tolerated in all patients. No evidence of acute haemorrhage, petechia, ischemia or oedema was observed in post-sonication SWAN T2*, Diffusion or FLAIR MRI sequences.

Conclusions

The BBB was safely opened by pulsed ultrasound using an implantable ultrasound device in patients with recurrent glioblastoma. Additional follow up and recruitment will be used to further evaluate the safety and potential efficacy of such an approach.

Acknowledgments

Work supported by CarThera and the Hospital Pitie Salpetriere.

O32 Investigation of temperature dependent changes in MR signal intensity, t1 and t2* in cortical bone

Henrik Odéen2, Bradley Bolster1, Eun Kee Jeong2, Dennis L. Parker2

1Siemens Healthcare, Salt Lake City, Utah, USA; 2Radiology, University of Utah, Salt Lake City, Utah, USA

Objectives

For MR guided focused ultrasound treatments in or close to bone, such as for transcranial applications focusing through the intact skull bone and treatments of bone metastases, significant heating can occur in the bone due to its high ultrasound absorption. MR imaging of bone is in general challenging due to the short T2 relaxation time of bone. For MR temperature imaging the short T2 also severely decreases the accuracy that can be achieved with the standard proton resonance frequency shift method. Instead researchers have investigated the temperature dependence of the MR signal intensity (SI) and T1 relaxation time for temperature monitoring (1–4). Miller (1) and Fielden (3) et al. showed that the SI from cortical bone decreases with increasing temperature using ultrashort echo time (UTE) pulse sequences, and Ramsay et al. (2) found that, contrary to what Miller and Fielden observed, the SI increases with increasing temperature using a short TE gradient recalled echo (GRE) pulse sequence. Han et al. (4) further showed that T1 increases with temperature, also using UTE.

In this work we investigate the temperature dependence of the SI (dSI/dT) and the T1 and T2* relaxation times (dT1/dT and dT2*/dT, respectively) using a 3D UTE pulse sequence to investigate which parameter has the highest sensitivity to temperature change.

Methods

All imaging was performed on a 3T MRI scanner (MAGNETOM PrismaFit, Siemens Healthcare, Erlangen, Germany) using a 3D UTE pulse sequence. The sequence utilizes radial, ramped sampling of k-space in 3D starting at the k-space centre after a 80 μs non-selective hard RF pulse, allowing TEs down to 50 μs. T2* was measured by performing an exponential fit to data acquired at TE = 50, 90, 130, 170, and 250 μs (other scan parameters are listed in Table 1). The 50 μs TE was also used for dSI/dT calculations. T1 was measured using the variable flip angle (VFA) method (5) with TR = 11 ms and FA = 8 and 36° (other scan parameters are listed in Table 1).

An approximately 4-cm long bovine femur (marrow and connective tissue removed) was placed in a phantom holder that allowed heated water to circulate around the bone, Fig. 27. One fiber optic probe measured the water temperature, and three probes were inserted in 1-mm diameter, 2-cm deep, holes drilled in the bone to measure the temperature of the bone. The water was warmed to ~22, 35, 50, and 65°C and data was collected when all 4 probes measured within 1°C. The whole setup was places in a 20-channel RF head coil for signal detection.

Results

Figure 28 shows 2D maps of SI and T1 and T2* relaxation times for the 4 different temperatures. Mean and standard error values from a 9x9 ROI close to each probe is shown in Fig. 29, together with calculated changes in %/°C. A decrease in SI of 0.3-0.5%/°C, and increases in T1 and T2* of 0.5-0.9%/°C and 0.6-0.9%/°C, respectively, was observed. Using the temperature dependent spoiled GRE signal equation (6) and the observed values for dT1/dT and dT2*/dT, dSI/dT can be closely predicted.

Conclusions

The decrease in SI is in accordance with previously published results by Miller and Fielden. The measured change in T1 using UTE agrees well with the 0.6%/°C reported by Han, although we measured higher absolute T1 values (~160-175 ms at 25°C, compared to ~115-125 ms as reported by Han).

The effect of T1 and T2* on SI are counter-acting each other (both increasing), which reduces the sensitivity of dSI/dT. This may suggest that dT1/dT and dT2*/dT are more suitable candidates for bone MR thermometry, and Fig. 29 also shows higher sensitivity for relaxation times that for SI. However, it should be noted that SI can be detected from a single image, whereas T1 and T2* measurements utilize two or more images, therefor resulting in longer scan times.

Future studies will acquire data with longer TEs (out to ~10 ms) to investigate the difference in temperature sensitivity for shorter and longer T2 components in the bone. These results will be compared to the increase in SI with temperature observed by Ramsay using a TE ≈ 1 ms. The temperature dependence of the long and short T2* components can be found by a multi-exponential fit. Flip angle mapping will also be implemented to improve the accuracy of the T1 measurements (7).
Table 1 (abstract O32).

Scan parameters

 

FOV (mm)

Res (mm)

TR (ms)

TE (ms)

FA (deg)

BW (Hz/px)

Scan time (s)

# views

UTE T1

160×160×160

1.0×1.0×1.0

11

0.05

8, 36

1008

167

15000

UTE T2*

160×160×160

1.0×1.0×1.0

6

0.05, 0.09, 0.13, 0.17, 0.25

15

1008

122

20000

FOV Field on view, Res Resolution, TR Repetition time, TE Echo time, FA Flip angle, BW Bandwidth (read out), # views – number of radial views/rays acquired

Fig. 27 (abstract O32).

Scan setup. A ~4 cm long bone sample was placed in a phantom holder that allowed water circulation to homogenously heat the bone. 4 fiber optic probes were used; 1 in the water and 3 in the bone sample.

Fig. 28 (abstract O32).

2D maps of SI and relaxation times for the 4 different temperatures (~22, 35, 50, and 65 °C), a SI, b T1, and c T2*

Fig. 29 (abstract O32).

Changes versus temperature for a SI, b T1, and c T2*. Mean and standard error value from a 9x9 voxel ROI close to each probe is shown

O33 Spatially-segmented MRI brain and water bath reconstruction for undersampled transcranial mr-guided focused ultrasound thermometry

Pooja Gaur1, Xue Feng2, Samuel Fielden2, Craig Meyer2, Beat Werner3, William Grissom1

1Vanderbilt University, Nashville, Tennessee, USA; 2University of Virginia, Charlottesville, Virginia, USA; 3University Children's Hospital, Zurich, Switzerland

Objectives

MR-guided focused ultrasound (MRgFUS) brain systems deliver targeted thermal energy into the brain using a hemispheric array of transducers that surround the head with an intervening water bath (Fig. 30a). During treatment the localized heating (hot spot) is measured from a change in image phase between baseline (pre-treatment) and dynamic (during treatment) images. Accelerating temperature mapping by undersampling k-space is desirable to increase spatiotemporal resolution and coverage, but is difficult to do with parallel imaging since coils must be placed outside the transducer, far away from the head. Multiple groups have instead developed accelerated temperature mapping methods that exploit temporal correlations between baseline and dynamic images [1, 2]. However, circulation of the water bath to cool the skull causes dynamic signal changes that are not captured by baseline images (Fig. 30b), which breaks those correlations and results in artefacts throughout the temperature maps. We propose a spatially-segmented approach for reconstructing temperature maps in brain MRgFUS, in which we separately estimate a water bath image without a baseline, and a temperature map in the brain with a baseline. The method can estimate artefact-free temperature maps from undersampled data during brain MRgFUS treatments using a single receive coil.

Methods

Our iterative approach alternates between updating the parameters of a k-space hybrid signal model which is fit in the brain region of the image [1], and a baseline-free estimate of the water bath image. The fitting of k-space hybrid brain model results in a phase drift-corrected brain image without the temperature phase shift and a sparse temperature phase shift map. An algorithm to fit the model is described in [1]. The water bath is reconstructed using a POCS algorithm that alternately enforces data consistency, consistency with a water bath support mask (brain and water bath masks are obtained from a baseline image), and sparsity in the Coiflet domain using soft thresholding [3]. Figure 30c illustrates the overall undersampled dynamic image model.

To test the method, a gel-filled human skull phantom was sonicated by an InSighTec ExAblate Neuro 4000 transcranial MRgFUS system (InSighTec Ltd, Haifa, Israel) while imaging with a GE 3T MR750 scanner (GE Healthcare, Waukeshaw, WI). 27 single-slice 2DFT gradient echo images were collected with the body coil and 28 x 28 x 0.3 cm3 field of view, 256 x 128 acquisition matrix, 30° flip angle, 13 ms TE, and 28 ms TR. Images and maps were reconstructed to a 128 x 128 matrix and retrospectively randomly undersampled by 2x, with full sampling over 22 central k-space lines. Temperature maps were reconstructed by fitting the k-space hybrid model to the entire image, or to the brain only with keyhole or POCS methods used to reconstruct the water bath image.

Results

Figure 30d shows the temperature reconstruction results. When the k-space hybrid model is fit to the entire image without distinguishing between the brain and water bath, phase artefacts obscure the hot spot in the reconstructed temperature map and (in this case) lead to an overestimation of the temperature rise in the sonicated region across image dynamics (RMSE across dynamics: 0.0121°C). Restricting the temperature reconstruction to within the brain, in combination with keyhole reconstruction of the water bath image (using the baseline image’s k-space to fill in missing k-space lines), produces temperature maps with lower errors in the hot spot but still large errors outside (RMSE across dynamics: 0.0039°C). The proposed k-space brain/POCS bath approach yields a more accurate estimate of the water bath image (not shown), resulting in much lower in-brain temperature artefacts (RMSE across dynamics: 0.0029°C).

Conclusions

Unpredictable water bath motion confounds model-based approaches to accelerated MR temperature mapping, resulting in large temperature artefacts due to aliased water bath signal. We demonstrated that a spatially-segmented reconstruction that applies a model-based reconstruction in the brain and a POCS reconstruction in the water bath can reconstruct temperature maps without undersampling artefacts at a moderate acceleration factor using a single receive coil. Future work will focus on integrating the approach with other accelerated temperature mapping methods [2] and extending it to non-Cartesian trajectories [4]. The method is compatible with multiple receive coils.

Acknowledgements

This work was supported by the Focused Ultrasound Foundation and NIBIB T32EB014841.

References

[1] Gaur P et al. Magn Reson Med 2015;73: pp. 1914–1925.

[2] Todd et al. Magn Reson Med. 2009;62: pp. 406–19.

[3] Lustig M et al. Magn Reson Med 2007;58: pp. 1182–1195.

[4] Fielden et al. Proc Intl Soc Mag Reson Med 23. 2015:1631.
Fig. 30 (abstract O33).

a During MRgFUS treatment, the patient’s head is immobilized in the transducer and circulating water bath. b The water bath signal varies significantly during a single focused ultrasound (FUS) sonication (arrow indicates sonication target). c In the proposed method, undersampled dynamic data are reconstructed using the k-space hybrid method in the brain and a POCS reconstruction in the bath. d Reconstructed temperature changes and maximum temperature errors in the brain with 2x undersampling. Temperature change averaged over the hot spot center is plotted at the bottom for each reconstruction. Circles on the x-axis indicate dynamics for which temperature maps are displayed above

O34 Efficient volumetric thermometry for MR-guided FUS brain treatment monitoring, using multiple-echo spirals and mixed update rates

Michael Marx, Pejman Ghanouni, Kim Butts Pauly

Radiology, Stanford University, Stanford, California, USA

Objectives

MR-guided focused ultrasound (MRgFUS) brain treatments are currently guided by one thermometry sequence: single-slice 2DFT MR thermometry. In this work, we divided treatment monitoring into two tasks, with different thermometry design goals for each, and developed sequences optimized for these goals: “Focal Spot Localization” and “Monitoring”. These sequences achieve greater imaging performance by utilizing multi-echo spiral thermometry, region-specific update rates, and MASTER slice interleaving.

Currently, focal spot targeting confirmation requires several low-power sonications to obtain high-resolution measurements in three dimensions. We developed a Focal Spot Localization sequence that obtains high-resolution measurements in-plane, at improved temperature precision compared to 2DFT, while also providing multiple-slices for through-plane characterization. This would reduce the number of sonications required, improving treatment time. Additionally, even lower-power sonications could be detected, improving patient safety.

During ablative treatment sonications, single-slice monitoring cannot detect through-plane shifts of the ultrasound focus, or unexpected out-of-plane heating. We developed a multi-rate thermometry Monitoring sequence that interleaves different sequences at different update rates to simultaneously achieve fast and precise focal monitoring, 3-dimensional focal spot measurement, and full brain monitoring. Fast and precise multi-slice monitoring of the focus ensures accurate thermal dose estimates for treatment feedback, while (slower) full-brain monitoring ensures patient safety.

Methods

All sequences were implemented using RTHawk (HeartVista, Menlo Park, CA) on a GE 3T 750 Signa scanner (GE Healthcare, Milwaukee, WI) equipped with the InSighTec Exablate Neuro (InSighTec, Haifa, Israel). All imaging was performed with the body coil, as is normally used with the Exablate system. Conventional 2DFT was implemented as a gold standard for comparison. The new sequences all use a 36 cm FOV to ensure that the transducer and water bath do not alias. Volunteer imaging was done with informed consent under IRB approval. Multi-frequency reconstruction was performed on all spiral data. Temperature uncertainty was measured as the per-voxel temporal standard deviation of temperature measurements, and averaged within manually segmented ROIs. Performance was also tested inside the transducer, using a gel phantom. Sequence parameters are compiled in Table 2.

The Focal Spot Localization sequence is a 5-slice 3-echo thermometry sequence, with doubled in-plane resolution as compared to conventional 2DFT (1.1x1.1 mm vs 1.09x2.18 mm), and acquisition time of 7 s. The Monitoring sequence interleaves 3 distinct sequences at different rates to monitor 29 total slices. The “Focus” is monitored using 3-slice 3-echo spiral imaging, for high-speed high-precision measurement of focal heating. Two adjacent slices, the “Boundary”, are monitored at half the temporal rate (also using 3-echo spiral) to fully characterize the focal spot. The remaining 24 slices, “Background”, were acquired using 8 blocks of 3-slice MASTER, with spiral readouts. Use of MASTER improves temperature uncertainty, compared to traditional slice interleaving, by increasing echo time. Limiting each MASTER block to 3 slices reduces inherent diffusion and motion-encoding artefacts. The three component sequences were interleaved such that Focus utilized 45% (15% per-slice) of the timeline, Boundary used 15% (7.5% per-slice), and Background the remaining 40% (1.7% per-slice).

Results

Figure 31a compares temperature uncertainty in vivo between 2DFT and Focal Spot Localization while Fig. 31b compares 2DFT with Monitoring for the same volunteer. Large images compare the centre slice, while stacks of images at the right show additional slices monitored by the new sequences. Figure 31c compares 2DFT and Monitoring uncertainty within the transducer. The new sequences obtained better uncertainty than 2DFT, with average values compiled in Table 2. Relative “Efficiency” is also listed in Table 2, which accounts for differences in speed and voxel volume. Each multi-echo spiral sequence is more than 150% as efficient as 2DFT. Background is 69% as efficient (but collects 3 slices per TR, for an effective 120% efficiency).

Conclusions

In this work, we have shown that multi-slice multi-echo spiral thermometry is an effective imaging approach for volumetric treatment monitoring. Improved imaging performance was successfully used to achieve imaging objectives for different aspects of ablative treatments. Focal spots may be localized faster and with reduced heating using the higher-resolution higher-precision Focal Spot Localization sequence. The mixed update rate Monitoring sequence successfully delivers high-speed high-precision monitoring of the targeted focus, while improving safety by simultaneously monitoring the full brain at a lower update rate. These sequences have also been validated within the transducer, to help ensure they will work in the clinical setting.
Table 2 (abstract O34).

Implemented Sequence Performance Parameters. tSNR efficiency is proportional to (δxyz*√(Tseq)*σT)−1 . For Monitoring - Background, median slice uncertainty reported

 

Speed (s)

FOV (cm)

Slices

Res (mm)

Slice Thickness

Uncertainty (phantom)

Uncertainty (in vivo)

in vivo tSNR Efficiency

2DFT

3.82

28

1

1.09×2.19

3 MM

0.48°C

0.66°C

100%

Focal Spot Localization

7.00

36

5

1.10×1.10

3 mm

0.16°C

0.38°C

253%

Monitoring - Focus

2.39

36

3

2.00×2.00

2 mm

0.15°C

0.41°C

182%

Monitoring - Boundary

4.78

36

2

2.00×2.00

2 mm

1.03°C

0.33°C

160%

Monitoring - Background

9.56

36

24

2.00×2.00

2 mm

0.15°C

0.54°C

69%

Fig. 31 (abstract O34).

Temperature uncertainty comparisons between 2DFT and a Focal Spot Localization, in vivo; b Monitoring, in vivo; c Monitoring, phantom within transducer. Large images compare centre slices, while stack of images at right show additional slices. Dotted outline in 1C delineates the phantom

O35 Towards MR-guided focused ultrasound treatments near metallic hardware

Hans Weber, Valentina Taviani, Kim Butts Pauly, Pejman Ghanouni, Brian Hargreaves

Radiology, Stanford University, Stanford, California, United States

Objectives

To demonstrate how MRI can be used in FUS treatments near metallic hardware for treatment planning, sonication monitoring and treatment assessment.

Methods

Using standard MRgFUS product sequences, we treated a 73-year-old patient with a metastasis in the right femur that was painful despite prior radiation and surgical stabilization with a metallic rod, demonstrating the inability of both echo-planar imaging (EPI) and gradient-recalled echo (GRE) imaging to be used for conventional proton resonance frequency (PRF) shift thermometry.

Based on the experience from this patient treatment, we have proposed an imaging strategy for MRgFUS near metallic hardware using multi-spectral imaging (MSI) techniques (Fig. 32a). Both MAVRIC-SL [Koch et al.; MRM 2011; pp. 65:71] and 2DMSI [Hargreaves et al.; ISMRM 2014, #615] are spin-echo-train-based imaging approaches that enable artefact-reduced imaging near metal. MAVRIC-SL compensates for distortions induced by field inhomogeneities by additional encoding along the slice dimension. Its ability to adjust the image contrast makes it a promising candidate both for treatment planning and assessment. 2DMSI limits the excitation to finite spectral and spatial regions (“frequency bins”) that can be imaged with minimal artefact quickly enough to be used for sonication monitoring.

We preliminarily tested the feasibility of MAVRIC-SL for treatment planning and assessment in a 65-year-old patient without metal hardware and with a metastasis in the right pelvis undergoing MRgFUS treatment for pain palliation. We used a GE 3T MRI system equipped with an InSightec ExAblate2000 FUS system. Proton density (PD) weighted MAVRIC-SL images were acquired prior to the treatment in addition to standard 2D fast-spin-echo (FSE) images. After treatment, T1-weighted MAVRIC-SL images were acquired before and after gadolinium injection in addition to standard 3D fast RF-spoiled GRE images with 2-point-Dixon fat suppression (LAVA-Flex, see Fig. 32c for details).

We tested the feasibility of 2DMSI to monitor a 25 s sonication in an acrylamide egg-white phantom containing the CoCr stem of a knee replacement and a pork loin sample containing a CoCr augment plate from the same hardware, each placed in a container filled with water. We acquired a time series of 10 single-slice 2DMSI images with a temporal resolution of 8 s/frame and the sonication starting after the third image. Each 2DMSI image was composed of 12 frequency bins ranging from −4.5 kHz to +5.4 kHz, and was acquired with TE = 30 ms, (bin) TR = 500 ms, 280 x 280 mm FOV, 5 mm slice thickness and 128 x 128 matrix size (effective number of phase encodings after half Fourier). To monitor the temperature-induced signal change, we subtracted the mean signal of the 3 baseline images from all 10 images voxel-by-voxel.

Results

Figure 32b shows the GRE and EPI image acquired in the patient with metallic hardware. With both techniques, the stabilized femur and its surrounding area are not visible due to distortions and signal dephasing. For comparison, the MAVRIC-SL image depicts the anatomy including the bone marrow surrounding the metallic rod.

Figure 32c presents the acquired images for treatment planning and assessment in the patient without metal. PD-weighted MAVRIC-SL achieves a contrast comparable that of the conventional FSE image and allows for localization of the metastasis in the lower part of the right pelvis. The T1-weighted MAVRIC-SL pre/post contrast difference image reveals the treatment area similar to the LAVA-Flex water difference image.

Figure 33 shows the thermometry results in the phantom and the tissue sample. In both cases, metal-induced field inhomogeneities of up to ± 4 kHz cause strong distortions and signal loss in the GRE image, whereas 2DMSI clearly depicts the area around the metal. In the phantom, the 2DMSI difference images yield a clear signal change at the focal spot next to the metallic stem and a noticeable change in signal at the focal spot next to the metallic plate in the tissue sample, despite a 70% reduction in SNR due to the lower water content. Averaging over the frames during sonication improves the localization of the focal spot.

Conclusions

We have presented initial results for our proposed imaging strategy for MRgFUS in the presence of metallic hardware.

MAVRIC-SL is an established technique for artefact-reduced imaging near metal. Here, we have shown that the image contrast can be adjusted to yield the relevant information for both planning and assessment of MRgFUS treatments. The reduced image resolution compared to the standard FSE protocol (to keep the scan duration at an acceptable length) did not noticeably reduce diagnostic image quality for the treating radiologist.

We have also demonstrated that 2DMSI enables the measurement of temperature-induced signal changes in close proximity to metallic hardware and thus in regions where conventional PRF shift thermometry fails. The bin-selective approach allows for a temporal resolution of 8 s/frame, which is sufficiently high to resolve the temperature evolution in sonications as short as 20 s, which is the typical duration in MRgFUS treatments. In case of less severe metal-induced artefacts, the number of frequency bins could be reduced to increase the temporal resolution. Averaging over time frames could facilitate the detection of the focal spot in lower SNR cases. For the given temporal resolution and the low temperature sensitivity of the T2 relaxation time of aqueous tissues, the measured signal change is expected to be highly dominated by the temperature sensitivity of the PD, whereas a higher temporal resolution is expected to increase the T1 weighting. While the latter provides higher temperature sensitivity and thus lowers the SNR requirements, PD weighting offers the benefit of a tissue-independent temperature mechanism that could facilitate quantitative thermometry.

In conclusion, the proposed imaging strategy has the potential to enable MRgFUS treatments near metallic hardware. The patient population at greater risk for cancers, and hence osseous metastases, overlaps with the elderly demographic more likely to have metallic hardware such as joint replacements. Further, orthopaedic bone stabilization is often used as a treatment for osseous metastases at risk for fracture. Overcoming these technical limitations is therefore important to allow the use of MRgFUS in a larger patient population.
Fig. 32 (abstract O35).

a Components of the proposed imaging strategy for MRgFUS near metallic hardware. b GRE and EPI images underlying conventional PRF thermometry and a MAVRIC-SL image, all acquired in a patient with a femur stabilized with a metallic rod. The dashed line in the x-ray image depicts the location of the MRI slices. c Comparison of both planning images and difference images (pre and post contrast injection) for treatment assessment, acquired in a patient without metallic hardware

Fig. 33 (abstract O35).

GRE and 2DMSI image and 2DMSI signal change during sonication in the phantom (a) and the ex vivo tissue sample (b). The 2DMSI difference images are cropped and masked to the dashed area. For the tissue sample, the 2DMSI difference image in the lower right shows the average over the three time frames during sonication

O36 Thermal monitoring of HIFU using thermal memory effect of phase-change nano droplet

Jun Tanaka, Kentaro Kikuchi, Ayumu Ishijima, Takashi Azuma, Kosuke Minamihata, Satoshi Yamaguchi, Teruyuki Nagamune, Ichiro Sakuma, Shu Takagi

The University of Tokyo, Tokyo, Japan

Objectives

Phase change nano droplets (PCND), whose diameters are 200–400 nm, are droplets of perfluorocarbon (PFC) covered with phospholipid layers. Since they can be vaporized by ultrasound and transformed to microbubbles, they will be utilized as ultrasound contrast agents and ultrasound therapy sensitizers. There are several types of PCND which are often used in research, whose internal compositions are different, such as perfluoropentane (PFP) and perfluorohexane (PFH).Its boiling point can be adjusted by changing the mixture ratio of PFP and PFH.

The lifetimes of generated microbubbles changed from PCND are different from type to type. In this study, the lifetime dependence to ambient temperature at the moment of its vaporization was investigated. If remaining efficiency depends on the ambient temperature, this drug has a potential to be used as indicator of thermal memory effects and temperature monitoring agents.

Methods

Experimental setup for ultrasound exposure and high-speed imaging are shown in Fig. 34.In this study, we used the two types of PCND, whose internal compositions were PFP and the mixture (MIX; PFP: PFH = 1:1), respectively. Their main differences are boiling points (PFP: 29°C, MIX: 40°C), and the boiling point of MIX is estimated by thermodynamic calculation.

For the observation of the vaporization of PCND, we used two-layer structure of polyacrylamide gel phantom (layer with PCND and layer without PCND), and set the layer with PCND at the focal point of the transducer. PCND were vaporized by ultrasound (5 MHz centre frequency, 5 cycle bursts, Peak Negative Pressure = 3.5 MPa), which were irradiated with an arbitrary ultrasound beam controller (Verasonics) and a linear array transducer (EUP-L73S, Hitachi Aloka Medical). Time-lapse behaviours of PCND through phase changes were recorded with the high-speed camera (HPV-1A, SHIMADZU), coupled with inverted microscope (NIKON Eclipse Ti-U). The ambient temperature (gel phantom temperature) conditions were controlled with a hot water bath and thermo plate (TOKAI HIT) on the stage of the microscope.

Results

First, we observed the vaporization of the two types of PCND at 37 °C. The high-speed images of vaporization at 37 °C are shown in Fig. 35.

As to MIX (B.P. = 40 °C), generated microbubbles disappeared soon (within 10 μs) after the ultrasound exposure. On the other hand, as to PFP (B.P.=29 °C ), generated microbubbles remained for a while (more than thirty seconds). Because the main difference of these two types of PCND is the boiling point, we assumed that temperature is the key factor, and tried to control the difference between the boiling point and the ambient temperature. Then, we did some vaporization experiment at various temperatures. Some of the high-speed images of vaporization at 26 °C, 48 °C are shown in Fig. 36.

As to PFP (B.P. = 29 °C), generated microbubbles disappeared soon after the sonication at 26 °C, although they remained for some time at 37 °C. As to MIX (B.P.=40 °C), generated microbubbles had long lifetimes at 48 °C, although they had very short lifetimes at 37 °C. It can be considered that the lifetimes of generated microbubbles are greatly affected by not only the internal composition, but also the ambient temperature. Besides, behaviours after vaporization at 26~48 °C are shown in Fig. 37. Whether generated microbubbles will remain or disappear was switched around the boiling point.

Conclusions

We found that the lifetime of microbubble highly depended on the ambient temperature at the moment of vaporization. This effect has a potential to be used as indicator of thermal memory effects and temperature monitoring agents.

Acknowledgements

Authors thanks to Dr. Kawabata and Mrs. Asami in Hitachi Central Research Laboratory.
Fig. 34 (abstract O36).

Experimental setup

Fig. 35 (abstract O36).

High Speed Images of Vaporization at 37 °C

Fig. 36 (abstract O36).

High speed images of vaporization at 26, 48 °C

Fig. 37 (abstract O36).

Behavior after vaporization at 26~48°C

O37 Localized blood brain barrier opening of the macaque brain using a high frequency 512-elements FUS transducer and ultrasound contrast agent

Mathieu D. Santin1,2, Laurent Marsac3,4, Guillaume Maimbourg5, Morgane Monfort2, Benoit Larrat4, Chantal François2, Stéphane Lehéricy1,2, Mickael Tanter6, Jean-Franҫois Aubry4

1Centre de NeuroImagerie de Recherche - CENIR, Paris, France; 2Inserm U 1127, CNRS UMR 7225, Sorbonne Universités, UPMC Univ Paris 06 UMR S 1127, Institut du Cerveau et de la Moelle épinière, ICM, Paris, France; 3Supersonic Imagine, Aix-en-Provence, France; 4Institut Langevin, CNRS, Paris, France; 5Institut Langevin, Université Denis Diderot, Paris, France; 6Institut Langevin, INSERM, Paris, France

Objectives

Blood Brain Barrier (BBB) protects the brain from most of the pathogen circulating within the bloodstream. Thus, the BBB renders a difficult brain penetration for most of the drugs used for brain therapy. Focused ultrasound has the ability to bypass the BBB in small regions considered as targets for various drugs such as chemotherapies. Being able to open the BBB in a restricted focal zone is a necessary step before using this method for therapies such as drug delivery. Here we developed an in-house primate-dedicated stereotactic frame mounted on a multielement focused ultrasound array to perform BBB opening, using combination of FUS and ultrasound contrast agent (UCA). Stereotactic images were loaded in a planning software that allows controlling the localization of the BBB opening. Sonications were real-time monitored using a passive cavitation detector (PCD) to detect signatures of stable or transient inertial cavitation originating from the UCA.

Methods

Experiments were conducted on an anesthetized macaque (Macaca fascicularis) each month during one year. Animal was anesthetized with a mixture of Ketamine (3mg/kg) and dexmedetomidine (15μg/kg) and the anaesthesia was maintained by an infusion of Alfaxan (Alfaxalone, 0.2mg/kg/min). The temperature of the animals was maintained at ~37°C using a heated water blanket. Animal physiology was monitored during the whole experiment. A homemade stereotactic frame holding the monkey head was affixed to a 512-element transducer resonating at 1 MHz (SuperSonic Imagine, France). Images were imported in a planning software in which all the positions of the head of the monkey in regard to the transducer were stored, along 6 axis of freedom (2 rotations and one translation for the transducer and one rotation and two translations for the head). Once the coordinate of the target was chosen in the stereotactic frame coordinates, the planning software allowed determining the position of the frame and the transducer. MRI was performed using a 3T Siemens Verio system (Siemens, Germany). Body coil was used for excitation and an 8-channel phased-array coil (Life Services LLC, USA) dedicated to primates was used for reception. T1 longitudinal relaxation time was obtained at baseline using an MP2RAGE sequence prior to BBB opening1. Ultrasound excitation consisted of a 0.6 MPa Peak Negative pressure (as estimated at focus in the brain) sinusoidal tone burst of 10 ms, with a pulse repetition frequency of 1 Hz during 120s. Excitation followed a bolus injection of 1.5 mL of the UCA (Sonovue, Bracco, Switzerland) and lasted for 2 minutes.

The backscattered signal from microbubbles during insonification was recorded using a wideband (−6dB bandwidth: 4.5 - 14.4 MHz) transducer (Imasonic, France) acting as a PCD. The PCD was fixed at the right temporal bone window of the monkey perpendicularly to the FUS beam. This minimizes signal contamination from the main excitation field.

A bolus of 1.5 mL of an MRI contrast agent (Dotarem, Guerbet, France) was injected 5 minutes after the end of ultrasound excitation. A second MP2RAGE dataset was obtained 10 minutes after contrast agent injection to monitor the localization of the BBB opening resulting in a T1 decrease in the region of interest (ROI) induced by the contrast agent.

A conventional clinical scan consisting on a T1-, T2- and T2*-weighted MRI and DTI was performed at the end of the experiments to assess clinical status of the animal.

Results

This study was conducted in a living macaque during one year. Figure 38 shows the apparatus in position with the monkey and inside MRI. After BBB opening, T1 decrease was obtained in the ROI defined by the planning software (example on Fig. 39), indicating that BBB was opened in the targeted ROI. T1 values were 1000 +/− 70 ms before and 662 +/ -31 ms after the HIFU procedure resulting in a ~33% decrease in T1. The size of the area of BBB opening was 3.2 mm in diameter, and 5.6 mm along the axis of the beam.

Figure 40 shows the typical spectra obtained before and during infusion of gas microbubbles. Level of harmonics increased significantly for third (+13.6 dB) and fourth harmonics (+21.6 dB), but no significant increase in broadband noise was detected, suggesting that no transient inertial cavitation occurred during the insonification. After the experiment, the animal recovered and no side effects were observed during the 3 weeks following each BBB opening procedure. At the end of all the BBB opening procedures, T1-, T2- and T2*-weighted scans and DTI did not show any evidence of tissue damage (oedema, haemorrhages or bleeding) induced by the ultrasound procedures (Fig. 41).

Conclusions

The procedure allowed successful repeated transient opening of the BBB in a small ROI in a living primate with no side effects. No signatures of transient inertial cavitation were detected during experiments. This suggests that this method is safe for the animal. This study will be replicated in other animals with the long-term objective of developing a system suitable for human applications.

Acknowledgements

This work was supported by the Bettencourt Schueller Foundation and the "Agence Nationale de la Recherche" under the program “Future Investments” with the reference ANR-10-EQPX-15.
Fig. 38 (abstract O37).

Picture showing the animal within the setup and inside the MRI

Fig. 39 (abstract O37).

a. Planning software and target ROI definition. b. T1 map before BBB opening. c. T1 map after BBB opening. d. T1 difference between both images, arrow indicates the BBB opening on the targeted ROI

Fig. 40 (abstract O37).

Representative spectra obtained with the PCD. Black color represents the spectrum before injection of UCA, grey colour represents the spectrum with UCA. More and higher harmonics are identified when UCA is present. Broadband noise level is not different with or without UCA

Fig. 41 (abstract O37).

a. Colour Coded fractional anisotropy. b. T1-weighted image. c. T2-weighted image d. T2*-weighted image. No signs of oedema, haemorrhages or bleeding could be seen 3 weeks after BBB opening

O38 Enhanced neurorestoration through triple treatment with focused-ultrasound facilitated delivery of the neurotrophic factor neurturin

Maria Eleni Karakatsani1, Gesthimani Samiotaki1, Shutao Wang1, Camilo Acosta1, Eliza R. Feinberg2, Elisa E. Konofagou1,3

1Biomedical Engineering, Columbia University, New York, New York, USA; 2Biological Sciences, Columbia University, New York, New York, USA; 3Radiology, Columbia University, New York, New York, USA

Objectives

Currently, the existing Central Nervous System (CNS) drug delivery techniques are confined to either targeted but invasive or to non-targeted and non-invasive methods. Focused Ultrasound (FUS) coupled with the systemic administration of microbubbles has been proven to open the Blood Brain Barrier (BBB) locally, transiently and non-invasively, thus facilitating the diffusion of neurotrophic factors. Neurturin, a member of the glial derived neurotrophic factors (GDNF) family has been demonstrated to have restorative effects on the depleted by Parkinson’s disease dopaminergic neurons (DA). Moreover, our group has shown the bioavailability and downstream signalling of Neurturin in wild type mice and the restorative effect in Parkinsonian mice. Despite the promising results, the potential of multiple treatments with Neurturin in reversing the disease phenotype is still to be determined. The aim of the current study was to investigate the neurorestorative effect of triple delivery sessions of the neurotrophic factor Neurturin in a Parkinsonian mouse model.

Methods

For this study, twelve wild type mice (12 months old) were infused with sub-acute dosages of MPTP causing apoptotic degeneration in the nigrostriatal pathway. After the stabilization of the lesions and the decontamination period, the entire cohort was sonicated on the left hemisphere (ipsilateral side) targeting twice the Caudate Putamen region (CPu), to cover the entire area, and once the Sabstantia Nigra region (SN). Half of the mice received an IV injection of 0.5mg Neurturin accounting for the treated group, FUS+/NTN+, while the rest constituted the control group, FUS+/NTN-. Magnetic resonance imaging (MRI) was performed after each sonication to verify the accuracy of the BBB opening in terms of targeting. The procedure was repeated once biweekly to a total of three treatments. Following the third treatment, the survival period lasted for 28 days letting the neurotrophic factor to develop its restorative effects. On the 29th day, the mice were sacrificed and coronally sectioned for tissue processing. The brain slices of both the SN and the CPu were stained for tyrosine hydroxylase positive cells (TH+) with a custom protocol. The stained slices were imaged to count the TH+ nerve cell nuclei on the SN while the axons and dendrites were quantified by a custom MATLAB algorithm. For each mouse the contralateral side was compared to the ipsilateral side to eliminate inter-animal variation in the number of nuclei and projections. A quantification algorithm was used to compute the percentage of the relative difference (RD) between the two hemispheres, i.e., RD = (Ipsilateral – Contralateral)*100%. The process was repeated for all slices that cover the entire SN region and averaged across the mice. The error of the technique was measured as the standard deviation from the mean.

Results

There was no significant difference in the number of neurons between the ipsi- and contralateral sides. This result was in accordance with our knowledge of Neurturin restoring impaired neurons and not regenerating them. The RD was found to be significantly higher for the FUS+/NTN+ compared to the FUS+/NTN- group. This significance strengthens with the negative percentage of the FUS+/NTN- group implying a possible sensitivity of the Parkinsonian brain in multiple sonications.

Conclusions

These findings indicate a potential of multiple treatments on the reversal of the Parkinsonian phenotype as is the first time significance is reported. To strengthen this argument a second cohort of 20 mice is currently undergoing the same treatment aiming to apply various other imaging and quantification techniques to investigate the restoration of the functionality of the previously depleted neurons. Nonetheless, the current findings are essential considering the therapeutic effect of multiple treatments with FUS enhanced drug delivery in patients.
Fig. 42 (abstract O38).

a Contralateral side (control) of TH+ stained neuronal cells. b Ipsilateral side (treated) of TH+ stained neuronal cells. c Statistical analysis of the averaged TH+ stained neuronal cells of the two groups

O39 Long term effects of single vs repeated low intensity pulsed focused ultrasound treatment with microbubbles

Zsofia I. Kovacs1, Tsang-Wei Tu1, Georgios Z. Papadakis1,2, William C. Reid2, Dima A. Hammoud2, Joseph A. Frank1,3

1Frank Laboratory, Radiology and Imaging Sciences, National Institute of Health, Bethesda, Maryland, United States; 2Center for Infectious Disease Imaging (CIDI), Radiology and Imaging Sciences, National Institute of Health, Bethesda, Maryland, United States; 3National Institute of Biomedical Imaging and Bioengineering, National Institute of Health, Bethesda, Maryland, United States

Objectives

One potential issue for using MR-guided pulsed Focused Ultrasound (pFUS) to open the blood brain barrier (BBB) is the lack of data on the long term effects. Safety determination in the brain have been limited to the MR characterization after repeated BBB opening that can be achieved without haemorrhage, oedema and behavioural changes in non-human primates [1,2]. We use multimodal imaging technics to characterize long term effects of pFUS + MB in the rat brain to evaluate the effects of repeated BBB opening by pFUS and microbubbles (MB) on morphology to the rat striatum and hippocampus as monitored by magnetic resonance imaging (MRI), positron emission tomography (PET) and histology over 12 weeks.

Methods

Female rats were divided into two groups and received either pFUS + MB (OptisonTM, GE Healthcare, Little Chalfont, UK) once or six times targeting the striatum and the contralateral hippocampus. 200 μl of MB were administered intravenously over 1 minute starting 30 sec before pFUS. Rats received 3 daily doses of 300 mg/kg 5-Bromo-2′-deoxy-uridine (BrdU, Sigma Aldrich, St. Louis, MO) intraperitoneally before sonication to label proliferating cells in vivo.

0.3 MPa acoustic pressure was applied in 10 ms burst length and 1% duty cycle (9 focal points, 120 sec/9 focal points - striatum, 120 sec/4 focal points - hippocampus) using a single-element spherical FUS transducer (centre frequency 589.636 kHz; focal number 0.8; aperture 7.5 cm; FUS Instruments, Toronto, Ontario, Canada). T2, T2* and Gd-enhanced T1-weighted images were obtained by 3.0 T MRI (Philips, Amsterdam, Netherlands), T2, T2*, diffusion tensor imaging (DTI) and chemical exchange saturation transfer (CEST) imaging was performed by 9.4 T MRI (Bruker, Billerica, MA). Parameters for DTI: 3D spin echo EPI; TR/TE 700 ms/37 ms; b-value 800 s/mm2 with 17 encoding directions; voxel size 200 μm, isotropic. Diffusion weighted images were corrected for B0 susceptibility induced EPI distortion, eddy current distortions, and motion distortion with b-matrix reorientation using Tortoise. Parameter for glucoCEST: 2D fast spin echo with (MT) and without (M0) magnetization transfer (MT) pulses (TR/TE 3.5 s/11.5 ms; in plane resolution: 200 μm, thickness: 0.8 mm; MT pulse: 3 μT, 1 s). The MT offset frequencies (Δω) were set from −2 kHz to +2 kHz with 100 Hz stepping to detect the proton metabolites of glucose (1.2 ppm, 2.1 ppm, 2.9 ppm). Fractional anisotropy (DTI-FA) and the asymmetry of magnetization transfer ratio (MTRasym) were derived for mapping structural injury and glucose metabolism.

Quantitative of glucose uptake was performed with FDG-PET (Siemens, Munich, Germany). Each rat received under anaesthesia (O2 3–4 L/min & Isoflurane at 3–3.5 4%) 0.7 – 1.1 mCi of 18F-FDG via tail vein injection and was allowed to regain consciousness for an uptake period of 30 minutes in total. They were anesthetized again and a PET/CT study was acquired using Siemens Inveon Multimodality scanner (Siemens Medical Solutions USA, Inc.). CT scan was performed for localization and attenuation correction purposes. PET images were reconstructed using OSEM3D/MAP algorithm, with Ramp projection filter, scattered corrected, 2 OSEM3D iterations, 18 MAP iterations, 128 × 128 image size, and approximately 0.5mm resolution at the centre of the field of view (FOV).

Animals were euthanized 6 or 12 weeks after the first pFUS treatment. Histological evaluation of brain and tracking of BrdU tagged cells was performed at different time points. Values were compared to baseline.

Results

Preliminary results showed contrast enhancement on T1-weighted MRI in rats receiving a single sonication, indicating BBB disruption in the striatum and the hippocampus. Gd-extravasation or T2 and T2* abnormalities were not seen in the brain 1 day post pFUS + MB at 9.4 T MRI. Hypointense regions appeared on T2* MRI 2 weeks after pFUS + MB (Figs. 43 and 44) consistently with microhemorrhage within the parenchyma that decreased in volumes by week 3. White matter fiber structure- and gray matter-abnormalities on DTI MRI were detected in regions with T2* abnormalities (Fig. 44) suggestive of increased astrogliosis (Fig. 44a) and transient axonal damage (Fig. 44b). GlucoCEST showed loss of contrast as early as 1 day post pFUS and these changes persisted up to week 3 (Fig. 44a).

Qualitative analysis of MRI and GlucoCEST as well as 18F-FDG uptake with PET showed no difference between the sonicated region and the contralateral hemisphere 6 weeks post sonication.

Conclusions

We have observed a complex graded molecular and cellular sterile inflammatory response in the brain up to 24 hrs after pFUS + MB. However, little is known about the long term effects in rats using advanced imaging techniques. The DTI data showed that pFUS caused a low degree of structural injury at the location of sonication. However, the decrease in glucose concentration revealed by glucoCEST indicated that the pFUS could cause hypo-metabolism in the brain even after 3 weeks post sonication. These preliminary results suggest the importance of long term monitor of the brain following low intensity pFUS + MB.

Further research investigations are in process to evaluate changes following multiple targeted treatments in the brain.

References

[1] Arvanitis, C. D., et al., Cavitation-enhanced nonthermal ablation in deep brain targets: feasibility in a large animal model. J Neurosurg:1–10, 2015.

[2] Downs, M. E., et al., Long-Term Safety of Repeated Blood–brain Barrier Opening via Focused Ultrasound with Microbubbles in Non-Human Primates Performing a Cognitive Task. PLoS One 10(5):e0125911, 2015.
Fig. 43 (abstract O39).

3.0T MR images of a rat brain show Gd-extravasation immediately after pFUS + MB and delayed haemorrhage at the sonicated tissue (left striatum and right hippocampus) associated with BBB opening

Fig. 44 (abstract O39).

DTI and glucoCEST (9.4T MRI) of the rat brain (baseline, 1 day, 2 weeks and 3 weeks post pFUS) show changes in the grey and white matter tract at both sonicated locations. 2 weeks after pFUS + MB increased fractional anisotropy on DTI suggests astrogliosis in the striatum (a) and axonal injury in the external capsule (b). Decreased signal intensity in glucoCEST indicates lower glucose metabolism at the site of the sonication (a)

O40 Low intensity pulsed focused ultrasound and microbubbles results in sterile inflammatory response in the rat brain

Zsofia i. Kovacs1, Saejeong Kim1, Neekita Jikaria1, Michele Bresler1, Farhan Qureshi1, Joseph A Frank1,2

Frank Laboratory, Radiology and Imaging Sciences, National Institute of Health, Bethesda, Maryland, USA; 2National Institute of Biomedical Imaging and Bioengineering, National Institute of Health, Bethesda, Maryland, USA

Objectives

Very little is known about the graded cellular and molecular responses in the brain following pulsed Focused Ultrasound (pFUS) coupled with microbubbles (MB) exposures being advocated to increase drug or gene delivery through the disruption of the blood brain barrier (BBBD). We investigated the proteomic changes in the brain in response to Pulsed Focused Ultrasound (pFUS) + intravenous (IV) ultrasound contrast agent MB associated with BBBD.

Methods

MRI-guided pFUS was performed at 0.3 MPa acoustic pressure, 10 ms burst length and 1% duty cycle (9 focal points, 120 sec/9 focal points) using a single-element spherical FUS transducer (centre frequency: 589.636 kHz; focal number: 0.8; active diameter: 7.5 cm; FUS Instruments, Toronto, Ontario, Canada). 200 μl of OptisonTM MB (GE Healthcare, Little Chalfont, UK) were administered IV over 1 minute starting 30 sec before pFUS. Gd-enhanced T1-weighted images were obtained with a 3.0 T MRI (Philips, Amsterdam, Netherlands). Quantitative protein and mRNA expression in the brain following pFUS + MB were analysed with Bio-Plex ProTM Assay (Bio-Rad Laboratories, Inc., CA), enzyme-linked immunosorbent assay (ELISA), Western blot, Real-Time PCR (RT-PCR) or immunofluorescent staining. Proteomics were normalized to sham and statistical analysis was performed by one-way ANOVA corrected for multiple comparisons. 2 fold increase in mRNA expression was determined as significant. Rats were injected with 8 mg/kg Rhodamine encapsulated magnetic polymers (Biopal Inc., Worcester, MA) 3 days prior to pFUS to tag splenic macrophages. No evidence of damage or microhaemorrhage was observed on histology.

Results

The results of harvested brains at various times post sonication were as follows:
  1. 1.

    pFUS + MB resulted in BBBD by T1wMRI and by histology (albumin staining) without evidence of microhaemorrhages;

     
  2. 2.

    pFUS + MB induced a rapid (within 5 minutes) increased expression of pro-inflammatory and anti-inflammatory cytokines, chemokines and trophic factors originating from components of the neurovascular unit lasting up to 24 hours post sonication;

     
  3. 3.

    Proteomic analysis revealed increased heat shock protein 70 (HSP70), tumour necrosis alpha (TNFa), interferon gamma (IFNg) and interleukin (IL) 1a, 1b, 2, 5, 6, 17 and 18 consistent with damage associated molecular patterns (DMAP) (Chen and Nunez 2010) and activation of nuclear factor kappa-light-chain-enhancer of activated B cells (NFkB) inflammatory pathways observed with sterile inflammatory response to injury or insult (Fig. 45);

     
  4. 4.

    RT-PCR demonstrated activation of inflammatory genes associated with NFkB pathway along with anti-apoptotic genes, immune cell chemoattractants, selectins and cell adhesion molecule (CAM);

     
  5. 5.

    Evidence of influx of fluorescent bead labelled splenic macrophages in the brain by day 6 post pFUS along with activated astrocytes and microglia consistent with mild injury to the parenchyma.

     

Conclusions

The temporal molecular response to pFUS + MB is indicative of sterile inflammatory response (Gadani, et al. 2015) in the parenchyma originating from neurovascular unit. The pattern of pro-inflammatory cytokines immediately after pFUS + MB exposure is consistent with sterile inflammation initiated by DAMP that are released in response to ischemia or trauma associated with sterile inflammation observed with mild trauma or ischemia [1,2]. Increases in monocyte chemoattractant protein (MCP-1), vascular endothelial growth factors (VEGF), stromal derived factor 1 (SDF-1), erythropoietin (EPO) and brain derived neurotropic factor (BDNF) are associated with BBBD as well as stimulating angiogenesis, neurogenesis and stem cells migration consistent with mild injury following pFUS + MB exposure to the brain. These results indicate that pFUS + MB rapidly effects to the cerebral vasculature as evident by BBBD in addition to the shockwave from MB collapse induces mild stress within various cellular elements in the neurovascular unit.

References

[1] Chen, G. Y., and G. Nunez, Sterile inflammation: sensing and reacting to damage. Nat Rev Immunol 10(12): pp. 826–37, 2010.

[2] Gadani, S. P., et al., Dealing with Danger in the CNS: The Response of the Immune System to Injury. Neuron 87(1): pp. 47–62, 2015.
Fig. 45 (abstract O40).

T2*wMRI at 3.0T performed on day 6 post pFUS + MB. Recruited Fluorescent (Fl) SPIO labelled splenic macrophages homing to sonicated brain. FlSPIO (orange), CD68 macrophages (green) and microglia Iba1 (red) are consistent with injury to brain

Fig. 46 (abstract O40).

Stackplot of significantly increased expression of both pro- & anti-inflammatory factors over time following pFUS + MB to the brain. Proteomic profile is observed with sterile inflammatory response in the brain (Left). Stackplot of significant mRNA >2 fold mRNA expression at various time points (0.5, 6 and 12 hrs) post pFUS + MB in the brain. Peak mRNA expression was observed at 6 hrs post sonication that persisted to 12 hrs (Right)

O41 Pulsed focused ultrasound wave reconstruction and mapping for blood–brain barrier opening

Jingjing Xia2, Po-Shiang Tsui1, Hao-Li Liu2

1Department of Medical Imaging and Radiological Sciences, Chang Gung University, Taoyuan, Taiwan; 2Department of Electrical Engineering, Chang Gung University, Taoyuan, Taiwan

Objectives

Burst-mode low-pressure focused ultrasound (FUS) has been shown to induce transient blood–brain barrier (BBB) opening, and has high potential for use in non-invasive and targeted delivery of therapeutic molecules into the brain. FUS-BBB opening requires imaging guidance mean during the intervention, yet current imaging technology only enables postoperative outcome confirmation. In this study, we propose an approach to visualize short-burst low-pressure focal beam distribution that allows to be applied in FUS-BBB opening intervention guidance.

Methods

An backscattered acoustic-wave reconstruction method based on synchronization between emission from focused ultrasound and receiving diagnostic ultrasound elements and passively beam formed processing were developed. FUS transducers with the frequency ranging from 0.5-2 MHz were employed, and a commercialized diagnostic ultrasound was synchronously integrated with short-burst FUS excitation (burst length ranging from 0.01 to 10 ms). In-vitro phantom experiments were conducted to evaluate the constructed mapping, to quantitatively analyse its performance, and to evaluate the focal beam reconstruction limit. In vivo experiments with prior MRI anatomical scans were conducted to verify the feasibility of guiding the transcranial FUS-BBB opening procedure as well as its BBB-opened accuracy and reliability on small animals.

Results

A focal beam can be successfully visualized at all FUS frequency exposures (0.5 – 2 MHz) without involvement of microbubbles or acoustic cavitation triggering. The detectable level of FUS exposure with 0.467 MPa 0.05 ms single-burst exposure was identified. The signal intensity (SI) of the reconstructions was linearly correlated with the FUS exposure both in-vitro and in-vivo (r2 both higher than 0.9).

Conclusions

We confirmed that focal beam pattern can be visualized and allow successful guidance of FUS-BBB opening in small animals, with the SI level of the reconstructed focal beam correlated with the success and level of BBB-opening. The proposed approach provides a feasible way to perform real-time and closed-loop control of FUS-based brain drug delivery.
Fig. 47 (abstract O41).

In vivo treatment example showing the use of selected FUS exposure level to perform FUS-BBB opening

O42 Early assessment of mr-guided fus thalamotomy using a diffusion weighted steady state MRI sequence in an in-vivo porcine model

Juan C. Plata1, Samuel Fielden2, Bragi Sveinsson1, Brian Hargreaves1, Craig Meyer2, Kim Butts Pauly1

1Radiology, Stanford University, Stanford, California, USA; 2Bioengineering, University of Virginia, Charlottesville, Virginia, USA

Objectives

Diffusion-weighted imaging has been used to evaluate tissues ablated tissues using MR-guided focused ultrasound (MRgFUS), including uterine fibroids, prostate, and brain tissue [1–4]. Quantitative studies in canine prostate found a 36% reduction in the apparent diffusion coefficient (ADC) after either high intensity ultrasound ablation or cryoablation of the prostate, despite differences in histology [4]. More recently we studied the evolution of the ADC decrease and found that the time-course for the onset of ADC decrease after ablation of the canine prostate was inversely correlated to the thermal dose achieved [5]. As a result, areas that saw high levels of thermal dose saw a more rapid irreversible decrease in ADC, making ADC an early marker for loss of tissue viability in the prostate. Diffusion-weighted EPI inside of the InSighTec ExAblate 4000 Neuro System following thalamotomy demonstrates poor image quality. As a result, T2-weighted imaging is the method of choice for lesion detection inside the transducer although it may not be the earliest marker for ablation. More recently, a double echo in steady state (DESS) pulse sequence has been proposed to monitor lesion development. DESS generates two images, the first echo is mostly a gradient echo (GRE), the second echo is mostly a spin echo with some diffusion weighting [7]. The purpose of this work was to investigate the time course of lesion contrast in a pig model of thalamotomy on a diffusion-weighted steady state sequence in comparison to T2-weighted FSE. In addition, we probe the thermal dose dependence of the contrast by evaluating thermal lesions of two different peak temperatures.

Methods

MRgFUS thalamotomy was performed in a porcine model (n=2) under MR thermometry guidance. In one lesion in one animal, image collection began approximately 40 minutes after a low peak temperature sonication Tpeak = 52°C in the thalamus. In a second animal, two high peak temperature lesions Tpeak = 60°C were created in the thalamus, and image collection began immediately. In all cases, double-echo in steady-state (DESS) and fast spin echo (FSE) T2-weighted imaging acquisitions were interleaved. The parameters for both sequences are summarized in Table 3. Contrast to surrounding tissue was computed for all time points using regions of interest determined after lesion detection.

Results

Example images demonstrating the lesion on DESS and on FSE after the lower peak temperature sonication are shown in Fig. 48. The lesion demonstrates higher conspicuity in DESS than FSE.

In the quantitative analysis, in all three lesions, DESS provided superior contrast to T2-weighted FSE images at the early time points (Fig. 49), which equilibrated at the later time points. This is presumably due to the mixed diffusion and T2 contrast for the steady state sequence. As edema increases, the steady state sequence loses its advantage over T2-weighted FSE.

Higher peak temperature lesions demonstrate a faster time-course than the lower peak temperature lesion, seen in Fig. 49. In fact, both high temperature sonication lesions were conspicuous within minutes on the DESS sequence.

Conclusions

DESS provides a higher contrast between the lesion and the surrounding healthy tissue early after treatment is completed. This will allow for an earlier treatment evaluation while the patient is still in the brain transducer. Future work will include an in-depth simulation analysis on how both the diffusion weighting and T2-weighting contribute to the lesion detection time-course in FSE and DESS.

Acknowledgements

PO1 CA159992, RO1 CA111981, FUS Foundation, UVA-Coulter Translational Research Partnership.

References

[1] Jacobs MA, et al. JMRI 2009; 29:649–656.

[2] MacDannold N, et al. Radiology 2006;240:263–272.

[3] Wintermark M, et al. AJNR 2014;35:891–896.

[4] Chen J, et al. MRM 2008;59:1365–72.

[5] Plata J, et al. 2015. Med Phys 2015:09:5130.

[6] Chen L, et al. JMRI 1999;10:146153.

[7] Plata J, et al. 2015. ISMRM:1652.
Table 3 (abstract O42).

Imaging Parameters

Pulse Sequence

Readout

Flip (°)

TE (ms)

TR (ms)

Miscellaneous

FSE

3D

90

76

2500

ETL=100

DESS

3D

15

4.5/37

20.8

Spoiler=10cyc/voxel

FSE Fast Spin Echo, DESS Double Echo Steady State, TE Echo Time, TR Reception Time, ETL Echo Train Length

Fig. 48 (abstract O42).

Lesion Detection Using Fast spin echo (FSE) and Double Echo Steady State (DESS) images. Lesion is not visible with FSE until 60 min while the lesion is detected using DESS at 46min

Fig. 49 (abstract O42).

Contrast between lesion and surrounding healthy tissue as a function of time for fast spin echo (FSE) and Double Echo Steady State (DESS) images. Contrast for DESS is initially higher for both treatments indicating DESS can serve as an early indicator of lesion formation in brain treatments

O43 Monitoring thermal lesion formation with a steady state mri sequence in an in-vivo porcine muscle model

Juan C. Plata1, Vasant A. Salgaonkar2, Matthew Adams2, Chris Diederich2

1Radiology, Stanford University, Stanford, California, USA; 2Radiation Oncology, University of California San Francisco, San Francisco, California, USA

Objectives

The apparent diffusion coefficient computed from diffusion weighted imaging has been shown to have a 36% signal drop following tissue destruction with high intensity ultrasound and cryoablation [1]. Since this MR contrast mechanism is endogenous, previous studies have looked at monitoring the apparent diffusion coefficient (ADC) [1], or temperature and ADC [2], during treatment in order to study the progression of the ADC signal to assess when tissue viability is lost. More recently, a double echo steady state (DESS) sequence (Fig. 50) has been proposed to achieve a similar goal faster and with improved registration between the temperature images and the lesion monitoring images [3]. Due to its sensitivity to ADC changes and its short TR, DESS could serve as a monitoring platform for both temperature and lesion formation. Since DESS is a steady-state sequence, there are multiple echo pathways that also contribute to the measured signal. The first echo, Echo1, is effectively a gradient echo dominated by free-induction decay from to the preceding RF pulse. Although there are multiple echo pathway contributions for the second echo, Echo2, it is mainly a spin echo of Echo1 of the previous repetition with an effective echo time of TE2=2TR-TE1. The purpose of this work was to assess whether DESS imaging can provide lesion formation information during thermal ablation of in vivo porcine muscle.

Methods

In order to assess the DESS sequence in vivo, interstitial ablations were performed within 40–45 kg farm pigs under MRI guidance. All animal experiments were reviewed and approved by our institution’s Administrative Panel on Laboratory Animal Care (APLAC). MRI compatible interstitial ultrasound applicators, consisting of a two element array of 1.5 mm x 10 mm tubular ultrasound transducers, each with independent power control, were used to generate thermal ablative lesions within in vivo porcine muscle. The ultrasound applicators, with integrated water-cooling, were inserted within 13g (2.4 mm OD) Celcon plastic implant catheters which were placed free-hand 10–12 cm deep into the inner thigh muscles. For each experiment, two distinct thermal lesions were planned. For the first trial, thermal ablation was performed using the distal transducer only; after imaging and allowing time for the cool-down, the ultrasound applicator was translated back 2 cm within the stationary catheter for repositioning prior to the second trial. A second thermal lesion was then produced using the proximal element. This translation and sequencing from distal to proximal element were used to isolate the thermal lesions. All ablative sonications (n=4) were 10 min in length, with approximately 5–6 W acoustic power at 7.45 MHz delivered. Applicator and sonication parameters were chosen to generate circumferential lesions with an extent of approximately 1cm away from the active transducer segment. A Double Echo Steady State (DESS) sequence was developed using HeartVista’s SpinBench and RTHawk platforms in order to monitor thermal lesion development. The phase of the first echo of the image was used to obtain temperature information using PRF thermometry while the magnitude of the second echo was used to detect lesion induced changes in signal intensity DESS magnitude images. At the end of the treatment DESS magnitude images were compared to contrast enhanced (CE-MRI) and gross histology.

Results

Temperature information was successfully obtained from the first echo, and the lesion was monitored using the second echo (Fig. 51). Lesion extent obtained from the magnitude of the second echo correlated well with CE-MRI and gross histology (Fig. 52).

Conclusions

Lesion formation was visible using the proposed DESS sequence, allowing for more direct lesion monitoring during treatment.

Acknowledgements

PO1 CA159992, RO1 CA111981, GE Healthcare.

References

[1] Chen J, et al. MRM 2008; 59: pp. 1365–72.

[2] Plata J, et al. 2015. Med Phys 2015:09:5130.

[3] Plata J, et al. 2015. ISTU Annual Meeting Abstract: 2173929.
Fig. 50 (abstract O43).

DESS Pulse Sequence Schematic. The Sequence consisted of two echoes a gradient echo (TE = 4 ms) and a spin echo (TE = 46 ms). A spoiler gradients of 60 cycles/cm or 6 cycles/voxel was used

Fig. 51 (abstract O43).

DESS Lesion Development. Magnitude images with a temperature change overlay are shown on the (row a) and magnitude images in the second echo (row b) are presented. Lesion development is clearly noticeable in Echo 2 and is present early during heating with contrast becoming stronger over time

Fig. 52 (abstract O43).

Lesion Extent Comparison. Magnitude images before and after contrast are compared to the magnitude image of the final Echo 2 image prior to contrast injection and gross histology

O44 T2-based temperature monitoring in bone marrow for mr-guided focused ultrasound

Eugene Ozhinsky, Matthew D. Bucknor, Viola Rieke

Radiology and Biomedical Imaging, University of California San Francisco, San Francisco, California, USA

Objectives

MR-guided focused ultrasound (MRgFUS) is a non-invasive technique for the treatment of painful bone metastases. Proton resonant frequency shift (PRF) thermometry is the standard method for monitoring temperature during MRgFUS interventions. It can precisely measure the changes in temperature in water-based tissues, but fails to detect temperature changes in bone and in tissues with high lipid content, such as bone marrow.

Current clinical protocols for bone treatments rely on measurement of the temperature change of adjacent muscle to estimate the temperature of the bone. This approach carries a significant risk of overtreatment in that more energy might be used than is needed to ablate the target. In fact, we observe in HIFU treatments of bone metastases that the highest temperature in soft tissue is only reached 10–15 seconds after the end of the sonication. Collateral treatment of the near-field soft tissues during MRgFUS increases the risk for muscle and vascular injury, which can result in significant perioperative or chronic pain.

Deeper penetration of the ablation through the cortical bone into the bone marrow or tumor is often desired for local control of osseous lesions. In the treatment of osteoid osteomas, complete ablation of the nidus is required for pain relief and to avoid recurrence, but the thickened cortical bone makes ultrasound penetration difficult. Therefore, temperature measurement within the bone is desirable.

Previous studies have shown a change in T2 of subcutaneous fat, red and yellow bone marrow in controlled calibration experiments and during treatments with focused ultrasound (Ozhinsky, et al. J Ther Ultrasound 2015; Baron et al. Magn Reson Med 2014). The goal of this study was to determine if T2 based thermometry could be used to monitor the temperature change in ex-vivo and in-vivo bone marrow during focused ultrasound ablation of intact bone.

Methods

All experiments were performed using an ExAblate 2100 system (InSighTec, Haifa, Israel) integrated with a 3.0 Tesla MR scanner (GE Healthcare, Waukesha, WI, USA). Bone marrow T2 was quantified with a double-echo fast spin-echo sequence with water suppression (TE = 35/186 ms, TR = 1500 ms, echo train length = 40, FOV = 32 cm, 128 x 128 matrix size, 10mm slice thickness, 15 sec/slice).

For ex-vivo validation, we performed MRgFUS ablation in an ex-vivo porcine femur (sonication: 20 sec, acoustic power: 30 W). The focus of the sonication was placed in the middle of the marrow, but due to the high ultrasound absorption of cortical bone most of the energy is absorbed in the cortical bone.

Focused ultrasound ablation was also performed in a swine model. All experimental procedures were done in accordance with NIH guidelines for humane handling of animals and received prior approval from the local Institutional Animal Care and Use Committee. Each of the three animals received 12–14 sonications on femur and ilium bones (acoustic power: 10-35W, duration: 20–40 sec). As in the ex vivo validation, the focus was placed in the middle of the marrow. At the end of the focused ultrasound, pre- and post-contrast 3D Fast SPGR images were acquired.

Results

Figure 53 shows the results of the ex-vivo experiment, where we measured a T2 elevation of 269 ms. Assuming the T2/temp coefficient of 7 ms/°C (Ozhinsky, et al. ISMRM 2014), this corresponds to a temperature rise of 38°. The ex-vivo experiment shows that it takes on the order of 15 minutes for the marrow to return to the baseline temperature.

Figure 54 shows the results of the in vivo experiment in a swine model. We measured a T2 rise of 231 ms within the bone marrow, which corresponds to temperature change of 33°C from baseline. The in-vivo experiment showed excellent correspondence between the area of T2 elevation in marrow during the ablation and the resulting non-enhancing area in the post-contrast images.

Conclusions

In this study we have demonstrated for the first time that T2-based thermometry can be used in vivo to measure the heating in the marrow during bone ablation. The ability to monitor the temperature within the bone marrow allowed visualization of the heat penetration into the bone, which is important for local lesion control and treatment of osteoid osteomas. Therefore, T2 based temperature mapping, in addition to PRF-based thermometry could be used to monitor heating during the bone focused ultrasound treatments and improve safety and efficacy of MRgFUS bone applications.
Fig. 53 (abstract O44).

T2 Measurement in ex-vivo bone marrow during and after the heating: a Localizer image showing the ultrasound transducer in the table; b T2 map during heating, showing the ROI; c Plot of T2 values within the ROI over time

Fig. 54 (abstract O44).

T2 Measurement in in-vivo bone marrow: a T2 map during ablation of a single sonication, showing the ROI; b post-contrast 3D Fast SPGR image after ablation (total of six sonications per location) c plot of T2 values within the ROI over time

O45 Tissue-mimicking thermochromic phantom for characterization of hifu devices, heating methods, and sonication parameters

Ari Partanen1,2, Andrew Mikhail2, Lauren Severance2, Ayele H. Negussie2, Bradford Wood2

1Clinical Science MR Therapy, Philips, Andover, Massachusetts, USA; 2Center for Interventional Oncology, Department of Radiology and Imaging Sciences, Clinical Center, National Institutes of Health, Bethesda, Maryland, USA

Objectives

Tissue mimicking phantoms (TMPs) are routinely used for calibration and quality assurance of medical devices including thermal therapy applicators prior to their use in clinic. TMPs are also used in thermal therapy research as alternatives to ex vivo soft tissues and organs as they possess several advantages including greater availability and shelf life, high uniformity, and customizability. The efficacy of thermal ablation therapies depends on several factors including targeting accuracy and temperature elevation in the treated tissue. Thus, an ideal TMP for thermal therapy applications should have the capacity to report ablated volumes and geometries as well as absolute temperatures. Magnetic Resonance Imaging -guided High Intensity Focused Ultrasound (MR-HIFU) is a therapeutic technique that can be used to precisely target and heat tissue non-invasively to induce thermal ablation or mild hyperthermia, among other applications. The objective of this study was to develop a novel, MR compatible tissue-mimicking thermochromic (TMTC) phantom for studying and characterizing HIFU devices, heating methods, and sonication parameters. Specifically, the intent of this work was to develop a phantom that reports on targeting accuracy, thermal energy deposition, and spatial heat distribution following HIFU. Additionally, the objective was to employ the TMTC phantom in characterization of two different HIFU devices, and to assess the temperatures and distribution of heating post-HIFU in relation to the treatment plan.

Methods

Polyacrylamide gel phantoms containing silica particles (1.0% w/v), bovine serum albumin (BSA, 3% w/v), and thermochromic ink (5.1% v/v, colour change temperature threshold of 60 °C) were produced. Both a preclinical Therapy Imaging Probe System (TIPS, Philips Research, Briarcliff Manor, NY) and a clinical MR-HIFU system (Sonalleve V2, Philips, Vantaa, Finland) were used for HIFU exposures targeted within the TMTC phantoms. The TIPS system contains an 8 element annular array with an 80 mm focal length, as well as a 2-axis motion control system to move between targets. The Sonalleve system contains a 256-element phased array transducer (focal length = 140 mm), as well as a motion control system with 5 degrees of freedom. HIFU exposure parameters for the TIPS were: frequency 1.0 MHz, acoustic power 30 W, with sonication durations of 60, 120, and 180 s. HIFU exposure parameters for the Sonalleve were: frequency 1.2 MHz, acoustic power 100 W, and duration 20–70 s, targeted to regions of 4–16 mm in diameter using electronic steering of the focal point. Together with the Sonalleve system, a clinical 1.5T MR scanner (Achieva, Philips Healthcare, Best, the Netherlands) was used for exposure planning and real-time thermometry utilizing the proton resonance frequency shift (PRFS) method. In addition, T2-weighted MR imaging and quantitative T2 mapping were performed to visualize and characterize thermal lesions within the TMTC phantoms after both Sonalleve and TIPS sonications. Post-MRI, HIFU-induced colour changes within the phantoms were identified, photographed, and compared to the sonication plan as well as to MRI T2 and temperature maps.

Results

Tissue-mimicking thermochromic phantoms were developed, produced, and validated for use in characterizing HIFU devices and sonication methods. HIFU thermal ablations (maximum temperature > 60 °C) resulted in permanent colour changes at targeted locations within the phantoms. These colour changes corresponded to maximum temperatures recorded using real-time MRI thermometry. A gradual colour change from yellow to magenta was visible between 40 °C and 64 °C. At temperatures above 64 °C, no further colour change was observed. In addition, heated volumes were visible on T2-weighted MRI and T2 maps as regions of permanent hypointensity and of lower T2 values, respectively, due to BSA coagulation. Increased HIFU energies and target diameters lead to greater colour change and a larger volume of colour change, respectively, as well as, correspondingly, to greater absolute change in T2 and a larger region of T2 change.

Conclusions

A tissue-mimicking thermochromic phantom was developed to assess the spatial targeting accuracy, maximum temperatures, and temperature uniformity of HIFU exposures. This TMTC phantom changes colour (over a range of temperatures that is relevant to ablative HIFU procedures) upon heating, allowing for quantitative measurements of absolute temperature and delineation of heated regions, and thus may be useful in HIFU device characterization, parameter optimization, quality assurance, and user training. TMTC phantoms can also provide volumetric temperature information in experiments where MRI-based real-time thermometry is not feasible, as the stepwise nature of the colour change associated with changes in temperature allow for assessment of temperature gradients within and at the periphery of the heated region.
Fig. 55 (abstract O45).

a HIFU target planning within a TMTC phantom performed on T1-weighted MR images using the Sonalleve therapy planning software. Each of the nine target locations is 12 mm in diameter. b HIFU thermal ablations produced temperatures above 60 °C as seen on real-time MRI thermometry. A coronal temperature map for a single sonication location is shown. c Intensity changes (due to BSA coagulation) on T2-weighted MRI with high spatial accuracy relative to the treatment plan. d HIFU thermal ablations resulted in permanent colour changes at the targeted locations within the TMTC phantom, correlating with the T2 changes and MR thermometry, and with high spatial accuracy relative to the treatment plan

O46 Reduced field-of-view MR thermometry in adipose tissues using zoomed apparent T2-mapping

Martijn de Greef2, Gerald Schubert1, Chrit Moonen2, Mario Ries2

1Philips Healthcare, Vantaa, Finland, 2Imaging Division, University Medical Center, Utrecht, Utrecht, Netherlands

Objectives

During HIFU ablation of abdominal and pelvic lesions, a balance needs to be established between ablation speed and heat accumulation in the pre-focal subcutaneous layers, resulting in a safe but time-efficient intervention. As the de facto standard form of magnetic resonance thermometry, which is based on the proton resonance frequency shift, is ill-suited for temperature monitoring in the adipose tissues, several alternatives have been proposed. Baron et al. (Baron, MRM, 2014) have demonstrated the feasibility of monitoring heating in adipose tissue layers using apparent T2 mapping based on dual echo fast spin echo imaging. A linear (5.2 ms/C) and reversible T2 - temperature dependency was shown reproducible over a relevant temperature range.

In this study, the strategy developed by Baron et al. was combined with reduced field-of-view (FOV) imaging using perpendicular selection gradients, in the literature referred to as zonally-magnified or local-look imaging (Mansfield, J. of Phys.,1988). Reducing the FOV in phase-encoding (PE) direction while preventing fold-over allows to improve the spatio-temporal resolution of apparent T2-based thermometry. This opens new possibilities such as near-field monitoring during sonication at relevant temporal resolutions using sequence interleaving but also high resolution inter-sonication monitoring of accumulative heating.

Methods

All imaging experiments were performed on a 1.5-T MR Scanner (Achieva, Philips Healthcare, Best, The Netherlands) using a research version of the Sonalleve HIFU Platform (Philips Healthcare, Vantaa, Finland) that is supplied with an integrated four-channel loop coil and a 16-channel back coil.

Two imaging experiments were performed in this study. Under normothermic conditions, a volunteer was imaged using a dual echo fast spin echo sequence, with a single slice placed at the location of the subcutaneous fat layer (coronal orientation, TE1/TE2: 11/140 ms, train length: 24 lines/excitation, FOV: 400 x 105 mm2, matrix size: 168 x 48). The purpose of these experiments was to demonstrate the feasibility of apparent T2-mapping using a reduced FOV in PE direction.

A next experiment aimed at demonstrating the feasibility of resolving a temperature gradient along the subcutaneous fat layer of a volunteer, comparing a cooled and slightly heated state. Heating (37C skin temperature) and cooling (14C skin temperature) were achieved using a water-filled cushion together with a circulation unit, which was placed outside of the MR room. Skin temperatures were monitored using a fibre optic probe (Luxtron, LumaSense, Santa Clara, CA), placed between skin and cushion. Imaging parameters were the following: single slice, sagittal orientation, TE1/TE2: 11.8/130 ms, train length: 20 lines/excitation, FOV: 200 x 52 mm2, matrix size: 200 x 40.

In both imaging experiments, the radio-frequency pulse generating the initial transverse magnetization was implemented as a fat-selective binomial pulse (1-2-1) and the corresponding selection gradient was oriented parallel and perpendicular to the imaging slice, respectively. As a consequence, the volume experiencing the spectrally non-selective refocussing pulses was limited to the subcutaneous tissue layers. This aims at preventing disturbance of the magnetization in the water compartment at the location of the focus.

Results

An overview of the results of the imaging experiment under normothermic conditions is shown in Fig. 56(a-c). Reduction of the FOV in phase-encoding direction (left-right in the image) was achieved without apparent fold-over artefacts.

Figure 56(d-f) shows the signal intensity at the first echo time (panel d), the signal intensity at the second echo time (panel e) and the difference in apparent T2 between the heated and cooled state as an overlay (panel f, background: intensity at the first echo time). A gradual change along the anterior-posterior direction (left-to-right in the image) is observed in the difference of the apparent T2.

Conclusions

Dual echo fast spin echo imaging in combination with reduced FOV imaging using perpendicular selection gradients was successfully shown to allow for apparent T2-mapping with no/minimal fold-over artefacts under normothermic conditions. Furthermore, comparing a cooled and slight heated state, a gradient in the apparent T2 difference could be resolved at high resolution (~1mm). Both examples show the potential of reduced FOV apparent T2-mapping, enabling near-field monitoring at improved spatio-temporal resolution. The potential interference in interleaved T2/PRFS imaging scenarios is currently under investigation.
Fig. 56 (abstract O46).

Overview of the result obtained in the two imaging experiments. The results obtained under normothermic conditions are shown in panels a-c, where panel a shows the signal magnitude at the first echo time, panel b the signal magnitude at the second echo time, and panel c the apparent T2. Panels d-f give an overview of the results obtained in a sagittal slice under cooled and slightly heated conditions. Panel d and e show again the signal magnitude at the first and second echo time, respectively, while panel f shows the difference in apparent T2 between the heated and cooled state

O47 Time-resolved in vivo measurements of FUS immunomodulation in a novel reporter mouse model of breast cancer

Megan E. Poorman1,2, Mary Dockery1, Vandiver Chaplin3,2, Stephanie O. Dudzinski1, Ryan Spears1, Charles Caskey4,2, Todd Giorgio1, William Grissom1, 2

1Biomedical Engineering, Vanderbilt University, Nashville, Tennessee, USA; 2Institute of Imaging Science, Vanderbilt University, Nashville, Tennessee, USA; 3Chemical and Physical Biology, Vanderbilt University, Nashville, Tennessee, USA; 4Radiology, Vanderbilt University, Nashville, Tennessee, USA

Objectives

Focused Ultrasound (FUS) therapy is a promising approach for treating cancerous lesions in the body. In addition to cell destruction, FUS hyperthermia has been shown to have immunomodulatory effects, increasing dendritic cell infiltration and activating the body's immune response to reduce metastases and future recurrence [1–2]. However, the development and understanding of FUS immunomodulation has been limited by an inability to characterize the immune response in vivo. This ability would enable optimal timing of ex vivo immunophenotyping, resulting in more efficient and more statistically powerful studies with fewer animals. Here we describe and validate a novel double-transgenic murine model of breast cancer that we have developed to meet this need.

Methods

A double transgenic PyNGL murine model was bred to express a nuclear factor-kappaB (NF-kB) reporter transgene (NGL) into the polyoma virus middle T oncogene model. This mouse exhibits spontaneous mammary tumour formation comparable to that of human breast cancer as well as a spatially-resolved, detectable change in bioluminescence with NF-kB activation, a key factor in immunomodulatory inflammation. By monitoring bioluminescence with In Vivo Imaging Systems (IVIS) imaging, tissue collection for immunophenotyping can be optimally timed. Experiments were performed to validate the mouse model when treated with FUS. Baseline luminescence maps were first obtained with IVIS prior to treatment. Subsequently, FUS thermal treatment was applied with either hyperthermia (CEM43 < 20) or ablative (CEM43 > 200) doses with a custom-designed MRgFUS system built in-house [3]. IVIS was used to monitor the change in luminescent inflammatory response every 12 hours after undergoing thermal therapy until the collection of tissues. This enabled the time of maximum immune response post-FUS treatment to be localized. Subsequent mice were sacrificed 48 hours post-treatment (the optimum time point based on the IVIS data) and analyzed with flow cytometry for infiltration of immune markers such as T cells (CD3,CD4,CD8), cytokines, and macrophages. Histology sections were taken of the skin, tumours, and spleen to assess cellular damage and composition (results not shown).

Results

For both treatment groups, a change NF-kB activation was observed with IVIS as soon as 24 hours post-treatment and reached a peak between 48 and 96 hours, consistent with the anticipated timeline for recruitment of inflammatory immune cells. Activation was spatially consistent with the area of MRgFUS treatment (Fig. 57a) and remained above baseline activation levels for the duration of luminescent imaging. The varied thermal doses were shown to have different effects on NF-kB activation - hyperthermia resulted in a consistent decrease in activation in the treated tumour (n=3) while ablation resulted in an increase in activation (n = 1) (Fig. 57b). Additional mouse studies are currently underway to confirm these results. Immunophenotyping revealed a large influx of T-cells 48 hours post-treatment in response to hyperthermia in comparison to the untreated control. Analysis of the ablated tumour tissue revealed no significant change in immune cell concentration compared to the control. (Figure 58). No superficial skin burns were observed on the treated mice in the area of applied MRgFUS treatment.

Conclusions

The use of a novel transgenic reporter mouse with spontaneous tumor generation enables spatiotemporally-resolved quantification of the immune response to FUS treatment in vivo. Cell analysis from excised tissue was supplemented by spatially-localized monitoring of the in vivo inflammatory immune response. Preliminary results showed that immune modulation measured by NF-kB activation depends on thermal dose. The increase in infiltrating T cells with hyperthermia, and lack thereof in the ablation case, suggests that the immune response may be more effectively activated by FUS treatment at a lower thermal dose. More experiments are ongoing to further explore the difference between thermal doses varying from hyperthermia to ablation as well as long term studies to investigate the effect of FUS-induced immune activation on metastases outside of the primary tumour.

Acknowledgments

This work was supported by DoD W81XWH-12-BCRP-IDEA, NIH RO1 CA111981, and a Vanderbilt University Discovery Grant.

References

[1] F Wu, Z Hu et al. Journal of Translational Medicine 5(34), 2007

[2] J Unga et al. Advanced drug delivery reviews 72:144–153, 2014

[3] M Poorman et al. ISMRM 2015 #1642
Fig. 57 (abstract O47).

IVIS data from MRgFUS treatment of mice. a Sample IVIS images of an NGL mouse (no tumours) showing a spatially-localized increase in NF-kB activity post-treatment consistent with where the MRgFUS was applied. b) Ratio of the average radiance of the FUS-treated tumour (HTT) to the ipsilateral tumour (ILT) over time plotted on a log scale. Two distinct dose-dependent treatment groups can be observed. Both treatments result in a large difference in NF-kB activity from baseline at 48 hours post-treatment

Fig. 58 (abstract O47).

A preliminary cellular analysis of treated tumours excised 48 hours post MRgFUS treatment and analyzed with flow cytometry. CD3+, CD4+, and CD8+ are cellular markers for T-cells that are involved in the in vivo immune response. A substantial difference in immune cell infiltration can be observed between the treatment groups

O48 Targeting tumour hypoxia with HIFU: a promising new adjuvant cancer therapy

Marcia M. Costa, Efthymia Papaevangelou, Anant Shah, Ian Rivens, Carol Box, Jeff Bamber, Gail ter Haar

The Institute of Cancer Research, Department of Physics, Sutton, United Kingdom

Objectives

Hypoxia is a common feature of radioresistant tumours, resulting in a decreased efficiency of radiotherapy (Harada, 2011). Several approaches have been proposed to overcome this limitation, including the adjuvant use of hyperthermia and chemotherapy. We propose the use of High Intensity Focused Ultrasound (HIFU) to target these hypoxic regions, which lack perfusion and so provide good targets for thermal ablation.

One of the challenges when targeting hypoxia is the accurate detection of poorly oxygenated regions, ideally non-invasively, for treatment planning. Several imaging techniques may be used in clinical practice for this purpose. Recently Photoacoustic (PA) imaging has been proposed for distinguishing between oxy- and deoxy-haemoglobin inside tumours. The technique is based on the generation of acoustic waves by the tissue after being exposed to short pulses of light. These waves can be detected using a transducer and the distribution of optical absorption in tissue is then reconstructed with US spatial resolution.

For this study, we used PA imaging to detect hypoxia in a radioresistant head and neck tumour model, implanted subcutaneously in mice. Selected hypoxic regions were targeted with HIFU, under US-guidance, using a dedicated small animal system, and the outcome of the treatment was evaluated histologically.

Methods

Twelve female NCr nude mice (6 weeks old) were subcutaneously injected with 5x105 CALR cells (head and neck tumour model) in the right flank. Tumours were measured up to 3 times a week, until they reached a volume of 200–300 mm3.

Animals were imaged using a MultiSpectral Optical Tomography (MSOT, iThera Medical) device with an excitation wavelength range of 660–1350 nm (details in Morscher, 2014). Coupling gel is applied to the tumour and surrounding area, and the animal is placed in a horizontal position in a holder under isoflurane anaesthesia. Furthermore, they are enveloped in a thin polyethylene membrane as it provides acoustic coupling, before being submerged in a water tank. Multiple transverse 2D slices of the tumour were acquired at 680, 700, 715, 730, 760, 770, 800, 850 and 900nm, in 0.5mm steps (from head to tail), and were reconstructed using interpolated model-matrix inversion (Rosenthal, 2010). The reconstructed data were multispectrally unmixed using a linear regression technique to identify the distribution of oxy and deoxygenated haemoglobin.

Animals were US imaged 24h after PA imaging (injectable anaesthesia: mix of hypnovel, fentanyl and medetomidine), using an E-cube scanner with a phased array transducer (f=12MHz). Animals were imaged in the same direction as in the PA experiment, with a 0.5mm step between each image acquisition. For the HIFU exposures, a pre-clinical VIFU2000 (Alpinion) system was used (single element spherical-focused transducer, 1.5MHz). Although hypoxic regions are not visible in B-mode images, they are generally distributed around areas of necrosis that can be identified by their hypoechogenicity. We compared these regions with those in PA images to define the target regions, which were exposed using the VIFU system at different acoustic power levels in order to define the thermal lesioning threshold for these tumours. Six animals were HIFU-exposed, at 31.1 +/− 3.1 W (N=2), 26.3 +/− 2.6 W (N=3) and 23.0 +/− 2.3 W (N=1), for 8 seconds, one exposure per tumour. In addition, a Passive Cavitation Detection (PCD) system consisting of a Precision Acoustics broadband sensor, 20 mm outer diameter, weakly focused co-aligned to the HIFU focal peak and connected to a data acquisition system (Spectrum MI.2031, 8-bit) via a 1.5MHz notch filter. A 20dB pre amplifier was used to record data at a rate of 12x1.5MHz in 5 of the treatments. Broadband activity between 3-9MHz was analysed to detect inertial cavitation. Animals were allowed to recover for 24h after HIFU, at which point an i.p. injection of pimonidazole (a marker for hypoxia) was given and they were sacrificed 45min later. Tumour samples were collected, snapped frozen in cardice and stored at −80°C for histological analysis, which included pimonidazole and H&E staining.

Results

PA imaging suggested that this tumour model develops a necrotic core surrounded by hypoxic areas, despite the well vascularised rim around the tumour, when the volumes reaches ~200 mm3, after ~14-26 days, (Fig. 59). This was confirmed by H&E and pimonidazole staining (Fig. 60), although the latter did not show as extensive hypoxic regions as did the PA imaging.

US imaging was able to reliably identify the largest necrotic regions, which we used to compare with PA images to identify target areas for treatment. Of the HIFU treatments performed, the lowest acoustic power was interrupted. The remaining exposures resulted in tissue lesioning, observed after H&E histological analysis, and were identified as further extensive areas of necrosis and haemorrhage. None of the exposures resulted in skin burns. One interesting effect was that treated tumours did not take up the pimonidazole dye, which suggests a significantly reduced tumour perfusion at 24h. Inertial cavitation was detected in both higher power exposures, but not at lower power exposures.

Conclusions

The hypoxia distribution observed with photoacoustic imaging correlated well with that expected from the literature on this tumour model (Box et al., 2013). For future studies, we intend to co-register the PA and US images for a more precise treatment plan and use a larger cohort of animals to validate the results obtained with both modalities, using the standard histological techniques - pimonidazole and H&E.

The pimonidazole results, intended to indicate the oxygen distribution, showed lack of penetration of pimonidazole, probably due to vascular occlusion, within 24 hours after HIFU treatment, whereas uptake in control animals processed at the same time was normal. Future studies will investigate this effect at different time points after treatment, with both post-exposure PA imaging and histological analysis, and investigating both hypoxia and perfusion staining. The effects of HIFU on vasculature have been studied before, but it is important to understand the time course of these effects in pre-clinical models as this may have a significant impact on the outcome of combined therapies, such as HIFU-chemotherapy and HIFU-radiotherapy.
Fig. 59 (abstract O48).

PA image of control tumour (green ellipse, 300mm3). Red: oxy-haemoglobin; blue: deoxy-haemoglobin

Fig. 60 (abstract O48).

Pimonidazole staining section, corresponding to the tumour in Fig. 59. Hypoxic areas are characterised by bright green regions, surrounding a (darker) necrotic core, as exemplified by the blue arrows

O49 Pulsed focused ultrasound stimulates the molecular responses necessary for stem cell homing through mechanical interactions with stretch-activated calcium channels

Scott R. Burks, Matthew Nagle, Ben Nguyen, Michele Bresler, Joseph A. Frank

Radiology and Imaging Sciences, NIH Clinical Center, Bethesda, Maryland, USA

Objectives

Stem cell therapies are promising regenerative medicine approaches. Pulsed focused ultrasound (pFUS) induces microenvironmental changes in normal and diseased tissues that can enhance local homing and efficacy of intravenously-infused mesenchymal stromal cells (MSC) and further improve disease outcomes. How pFUS interacts with tissues to produce the necessary molecular changes is unclear. Mouse muscle tissue was sonicated at increasing powers while passive cavitation was measured. Sonicated tissue was harvested for cyclooxygenase-2 (COX2) expression to correlate with physical ultrasound effects, as COX2 expression is critical for molecular signalling cascades that induce MSC homing. Lastly, mice were given inhibitors of mechanosensitive channels prior to pFUS to investigate their role in propagating physiological effects of pFUS.

Methods

C3H mice were treated with pFUS to the hamstring using a VIFU 2000 system. Under ultrasound imaging guidance, pFUS was delivered at 1 MHz, 5 Hz pulse repetition frequency, 5% duty cycle, and varying transducer output powers (ranging from 10–80 W). Passive cavitation detection was performed with a hydrophone. Mice were euthanized 16 hr post-pFUS and harvested muscle was homogenized and analyzed for COX2 expression by ELISA. For drug studies, mice were given GdCl3 (0.04 mmoles/kg) or ruthenium red (0.01 mmoles/kg) by intravenous injection at the beginning of sonications. Statistical comparisons were performed by one-way analysis of variance (ANOVA) using Bonferroni post-hoc tests with p values <0.05 considered significant.

Results

Statistically significant increases in COX2 expression were measured at 20, 40, 60, and 80 W compared to untreated muscle. COX2 expression measured after sonications at 60 and 80 W were significantly greater than COX2 expression after sonications at 20 or 40 W. Statistically significant increases in cavitation were not observed at 20 or 40 W, but were observed at 60 and 80 W. Increases in COX2 expression in were blocked at 20 or 40 W when either Gd or ruthenium red were administered during sonication. Forty watts was previously the maximum power that produced the necessary molecular changes without inducing tissue damage and therefore, was the maximum power investigated in this study.

Conclusions

Mechanical influences from pFUS drive molecular changes in tissue that are critical to stem cell homing processes. We have previously determined that COX2 expression is an acceptable proxy for molecular outcomes. At lower powers (20 and 40 W), cavitation from the sonications is not detectable, suggesting that cavitation-independent mechanical forces (i.e., acoustic radiation forces) drive COX2 expression. At higher powers (60 and 80 W), cavitation is detectable and COX2 expression is elevated compared to sonications at 20 and 40 W. It is unclear whether the cavitation detected at these powers drives the additional COX2 expression, or if it is the result of increased acoustic radiation forces at those powers. Regardless, 40 W was maximum power we previously determined not to cause detectable tissue damage and from the point of view of regenerative medicine, would be the maximum power used for those applications. At these powers, COX2 increases were blocked by Gd, a non-specific mechanostretch receptor blocker, and ruthenium red, a more specific blocker of transient receptor potential (TRP) channels. At the powers used for regenerative medicine, the acoustic radiation forces from pFUS activate TRP-class calcium channels to initiate the molecular cascade that necessary to induce stem cell homing.

O50 Pulsed focused ultrasound increases renal expression of interferon-gamma to enhance potency of mesenchymal stem cells and further improve acute kidney injury outcomes

Scott R. Burks, Matthew Nagle, Ben Nguyen, Michele Bresler, Saejeong Kim, Blerta Milo, Joseph A. Frank

Radiology and Imaging Sciences, NIH Clinical Center, Bethesda, Maryland, USA

Objectives

Pulsed focused ultrasound (pFUS) enhances homing of IV-infused mesenchymal stem cells (MSC) to murine kidneys during cisplatin (CIS)-induced acute kidney injury (AKI). pFUS acts as a neo-adjuvant to MSC therapy and the combination leads to better AKI outcomes (renal function and survival) than MSC alone. In wild-type mice, nearly twice as many MSC home to diseased kidneys following pFUS, but >10 times as much interleukin (IL)-10 is produced by MSC that home to pFUS-treated kidneys. This suggests that pFUS sonications modify the renal microenvironment to increase potency of MSC that home to sonicated kidneys. Interferon-g (IFNg) has long been known to increase MSC potency and has been shown to be upregulated in kidneys after pFUS. This study investigates the role of IFNg released by kidneys in response to pFUS improving the therapeutic efficacy of IV-infused MSC.

Methods

IFNg knockout (KO) mice received CIS (15 mg/kg ip), kidney pFUS (4 MPa; 5% duty cycle) and/or MSC (106 human MSC). Intravenous MSC injections were performed 3–4 hr post-pFUS. Groups included mice that had AKI only, AKI+pFUS, AKI+MSC, AKI+pFUS+MSC, and normal mice. Mice received CIS on Day (D) 0 and pFUS/MSC on D1. Some mice were euthanized on D2 and kidneys were harvested for molecular analyses. Other mice were euthanized on D4 to measure renal function (blood urea nitrogen [BUN]; serum creatinine [SCr]). Statistical comparisons were performed by one-way analysis of variance (ANOVA) using Bonferroni post-hoc tests with p values <0.05 considered significant.

Results

Following pFUS to the kidneys of IFNg KO mice, MSC homing to sonicated kidneys was enhanced ~2 fold compared to untreated contralateral controls. However, increased MSC homing did not lead to improved AKI outcomes compared to mice that received MSC injections alone. Levels of BUN and SCr, as well as expression of kidney injury molecule 1 (KIM1), were all significantly reduced by MSC treatment alone, but not further reduced by combination pFUS/MSC treatment like was previously observed in wild-type mice. Furthermore, significantly greater human IL-10 (IL-10 produced by MSC) was not observed in the pFUS+MSC group compared to mice that received MSCs alone.

Conclusions

pFUS creates a molecular zip code in AKI kidneys that enhance homing permeability and retention (EHPR) of infused MSC. While MSC infusions alone improve AKI to some extent in IFNg KO mice, the combination of pFUS+MSCs does not yield further improvements in disease outcomes like it did in wild-type mice. This demonstrates the pFUS-independent mechanism of AKI repair by MSCs does not require renal IFNg, but that the pFUS-dependent mechanism of improved repair/recovery does. It is likely that the IFNg released by pFUS is not solely responsible for potentiation of MSCs, but rather works in concert with a number of other immunological signaling molecules to achieve increased potency. However, IFNg appears to be the critical link for pFUS to function as a neo-adjuvant to MSC therapy in AKI as it is released and/or produced following pFUS. While functional outcomes correlate with lack of IL-10 production by MSCs in the IFNg KO mice, further studies will be necessary to elucidate its role in AKI recovery. These data provide molecular insight to justify using pFUS as a modality to improve MSC therapy during AKI, which often has limited therapeutic options clinically.

O51 Shear-wave manipulation for tracking high-intensity focused ultrasound (HIFU)

Nhan M. Le1, Shaozhen Song3, Kanheng Zhou1, Ghulam Nabi2, Zhihong Huang1

1Mechanical Engineering, University of Dundee, Dundee, Angus, United Kingdom; 2University of Dundee, Dundee, United Kingdom; 3University of Washington, Seattle, Washington, USA

Objectives

Focused ultrasound can induce shear-wave when certain requirements of acoustic power are met, e.g. pulse length, amplitude, frequency. Our group recently found that the characteristics of HIFU-induced shear-wave (HiSW) can be manipulated by changing different acoustic pulse-length and amplitude. By manipulating the HIFU pulse, we could control the wavelength and displacement of the HiSW. Our objective is to demonstrate that HiSW can be manipulated into either i) small and sharp propagating wave, or ii) long wavelength and large displacement propagating wave.

Methods

Optical Coherence Tomography with Phase-sensitive technique was utilized for shear-wave imaging. Phase-sensitive Optical Coherence Tomography performs a 256 (time axis) x A-scanlines (depth axis) over the period of 2 milliseconds at the same location for each HIFU pulse. This PhS-OCT-scan repeats in 256 different locations (width axis), forming a complete B-scan dataset over time (3D-dataset). The camera runs at 46200 kHz A-scan rate, exposure 17.4 μs, sensitivity 450 e/count.

Ex-vivo porcine skin is embedded inside 2%-agar phantom to ensure good contact with HIFU transducer. The HIFU transducer (2.09 MHz, 20 mm diameter, 13 mm focal length) is placed at the bottom of the sample. The scanning plane captures the top surface of the ex-vivo sample, defined by the centre point of the HIFU focus, and the axial- and lateral-direction of the HIFU beam.

HIFU-induced shear-wave (HiSW) is captured in reducing number of cycles per pulse, from 100 cycles/pulse to 20 cycles/pulse. The captured image is then processed offline for quantitative analysis of the HiSW, regarding the wavenumber of the HiSW signal.

Results

The displacement of HiSW correlates well with the reduction of HIFU cycles per pulse. Lowering HIFU cycles-per-pulse number would reduce HiSW displacement. A reduction of 80% HIFU cycles-per-pulse number (100 vs. 20 cycles/pulse) results in a reduction of less than 50% HiSW displacement (approximately 140 nm vs. 90 nm). Reducing the HIFU cycles-per-pulse parameter also leads to a sharper HiSW, regarding the peak wavenumber of HiSW (refer to Fig. 62, peak wavenumber of 4.2 mm- 1 in 100 cycles per pulse, as opposed to 3.7 mm−1 in 20 cycles per pulse). However, the sharper HiSW is greatly burdened by the relatively short propagation, as the high-frequency components quickly attenuate over travelling distance.

Conclusions

Understanding the response of HIFU-induced shear-wave (HiSW) under different acoustic settings, we can adapt HIFU into both diagnostic and treatment regime. In particular, our experimental setup would benefit the diagnostic and treatment of skin cancer at the same time. PhS-OCT can recover elasticity information from HiSW speed map. The acoustic power output is relatively low, with Isata measuring 23.6 mW.cm−2 maximum (100 cycles per pulse, total acoustic power output of 3.72 W), which is suitable for diagnostic purposes. By manipulating the acoustic settings, we could either a) induce a sharp HiSW for tracking purposes, or b) induce a strong HiSW with long propagating distance for diagnostic purposes, i.e. elasticity measurement.
Fig. 61 (abstract O51).

Displacement of HIFU-induced shear-wave at different time frame (after the first acoustic output pulse), with constant acoustic power output of 3.72 W, using a) 100 cycles per pulse, b) 20 cycles per pulse

Fig. 62 (abstract O51).

Wavenumber comparison between HIFU-induced shear-wave using 20 cycles per pulse versus 100 cycles per pulse, at constant time frame (85.4 μs after acoustic output pulse) and constant acoustic power of 3.72 W

O52 The twin piezo motor: low frequency miniature transducer

Shmuel Ben-Ezra1, Shani Rosen2

1Action-Physics, Pardes-Hanna, Israel; 2CByond ltd., Nesher, Israel

Objectives

Traditional transducer designs require a half-wavelength resonator, posing a lower bound for the size of the transducer. The Langevin type transducers are common for frequencies of about 30–60 kHz, comprising a stack of piezoelectric rings pressed to a metal resonator. Their size is about 50–100 mm, depending on frequency and speed of sound in the resonator material. This type of transducers is employed in ultrasonic cleaning baths, ultrasonic dental scalers, ultrasonic scalpel and in many other applications. The low frequency and the relatively large displacement-amplitude provided by the transducer make it ideal to generate cavitation, which may be useful, in turn, for the fragmentation of solid structures, plaque removal, tearing and cutting biological material etc.

In this paper we describe the development of the Twin Piezo Motor: a miniature low frequency, large displacement-amplitude transducer of a new design, in which the resonating element is a metallic, beam shaped tip, vibrating in a transverse (bending) mode. Geometry of the tip may be adjusted for a resonance in the required frequency range, while preserving small footprint.

The Twin Piezo Motor is assumed to be capable of generating cavitation in applications where space is heavily restricted. One possible application is Ultrasonic Lithotripsy (USL) - breaking kidney stones by cavitation. The required solution should pass through the urinary tract over a ureteroscope or a catheter, achieving full contact or getting very close (<2 mm) to the target stone. Upon activation, the transducer tip is assumed to generate a cavitation cloud on the adjacent stone surface, causing its fragmentation into small enough pieces.

Methods

We developed a series of transducer models, based on theory and finite element simulations (COMSOL Inc.). Some of the models had actually been built and tested. We used electric impedance analysis (LCR 3532–50, HIOKI Inc.) to locate resonances of the transducer; by comparison with simulation results we could identify the mode of vibration. Fast camera (Phantom V7.3 Turbo, Vision Research Inc.) equipped with a 200 mm lens (AF Micro-Nikkor 200mm f/4D IF-ED, Nikon Inc.) was used to record tip movement and cavitation dynamics in water. The setup enabled recording at frame rate of about 300,000 frame per second. Back illumination was used to enhance contrast. For driving the transducers we used a function generator (AFG 1022, Tektronix Inc.) and power amplifier (2100L, E&I ltd.), with custom transformers for impedance matching.

Results

We started by investigating the dental scaler (Selector U2 Plus, Apoza ltd.); we showed that it generates intense cavitation in water. Also, it can fragment a piece of chalk in water. A movie of 300,000 frames per second was produced, showing the dynamics of a cavitation cloud on a water-solid interface at a distance of 1 mm from the vibrating tip. Another interesting movie demonstrated the generation of mist by ultrasonic energy.

The transducer of the dental scaler is of the Langevin type; it is quite large, located inside the hand-piece, and the acoustic vibrations propagate along a shaft from the transducer to the active tip. We looked for a method to generate similar tip vibrations, producing the same results of cavitation, but with much smaller device.

The concept of the Twin Piezo Motor was developed, where the resonance is determined by transverse vibrations of the tip. Two piezoelectric bars of opposite polarity serve as piezoelectric engine such that when the left bar elongates, the right one contracts and vise versa. This combined motion is assumed to excite the transverse mode in the tip. Finite element simulations supported the design of the transducer, and first few samples were built, having overall length of 19 mm and width of 7 mm. The existence of the tip resonance was verified by impedance analysis: a minimum appeared in the anticipated frequency, and it was invariant under structural variations of the transducer.

The large amplitude vibration of the Twin Piezo Motor occurs at selective frequency and was demonstrated visually by the generation of mist. A nice movie at frame rate of 6688 frames per second was recorded.

Conclusions

The design of the Twin Piezo Motor is based on the resonance of the tip in transverse (bending) mode, with 2 piezoelectric bars at opposite polarities serving to generate the vibration. We demonstrated the resonative vibration of the tip and the generation of mist by tip vibrations. When the vibratory tip is in full contact with a wet stone, the device exhibits some grinding capabilities, reducing the size of the stone. Direct evidence for cavitation in water is still missing. The simple assembly of the transducer is done by one central bolt holding the components together. The first samples that we have built are too big; they have to be further diminished in a factor of 2 or 3 in order to pass through urinary tract.
Fig. 63 (abstract O52).

Schematics of the Twin Piezo Motor: 2 piezoelectric bars at opposite polarity are pressed by a central bolt between the back mass and the tip. The tip is designed to be in resonance in bending mode at the required frequency

Fig. 64 (abstract O52).

COMSOL simulation of the Twin Piezo Motor: The resonance frequency is determined by tip transversal vibrations

Fig. 65 (abstract O52).

First prototype of the Twin Piezo Motor. The vibratory tip is located close to the surface of an artificial stone

O53 Measurement of sonication duration for ablation of tumour in liver using trans-fusimo treatment system by using fiber-optic hydrophone

Senay Mihcin1, Jan Strehlow2, Ioannis Karakitsios1, Nhan Le1, Michael Schwenke2, Daniel Demedts2, Paul Prentice1, Sabrina Haase2, Tobias Preusser2, Andreas Melzer1

1School of Medicine, IMSaT, Dundee, United Kingdom; 2Fraunhofer, Mevis, Bremen, Germany

Objectives

The application of Focussed Ultrasound (FUS) in upper abdominal organs is particularly challenging due to complexity of breathing motion, a multitude of risk structures and possible occlusions through the rib cage. TRANS-FUSIMO Treatment System (TTS) is a newly developed software (MEVIS, Fraunhofer, Bremen) enabling Magnetic Resonance (MR) guided FUS (MRgFUS) in upper abdominal organs. MR organ motion tracking data is used for a model-based motion compensation while monitoring the temperature [1]. Due to the complexity of the system, TTS demands thorough validation before its use in animal trials. One important system parameter in this evaluation of FUS is the duration of sonication. The delay when starting a sonication, the deviation from the planned sonication time, and the delay in the system after a sonication is stopped, needs to be measured accurately to quantify system performance.

Methods

The evaluated version of the TTS uses the software interface of Signa 1.5 T MR Scanner (GE Healthcare, UK ) and the transducer of Conformal Bone System 2100 (CBS) (INSIGHTEC, Israel) . On the CBS transducer system, steerable sonications are realized by so called subsonications that have an individual focus position, and duration. Subsonications are organized in sonication banks and the active subsonication can be switched during sonication rapidly (2ms). Starting a sonication on the CBS, however has a considerable delay of approximately 2–3 seconds. To enable precise control of the sonication time, TTS employs a sonication strategy that exploits the short switching times between subsonications. For example, to start a static sonication, it builds a sonication bank with two subsonications. The first is a subsonication with very low power (0.001 W) and a long duration. It is the default subsonication and used to start the sonication. In the TTS, this step is called arming the transducer. The second subsonication has the actually prescribed sonication focus position and power information. It is activated immediately after the user choses to start the sonication via the TTS execute command. For safety reasons, the second sonication’s duration is limited to 250 ms and it is actively looped by the software until the prescribed duration of the sonication has passed, or until the user stops the sonication via a stop command (Fig. 66).

The sonication is monitored via single shot EPI MR Sequence of 512 phases, with TE: 26.4, TR: 100, flip angle: 40, freq phase: 128 x 96 parameters on the MR machine. To synchronise the MR to the TTS, the MR is configured to start monitoring after a TTL-Trigger pulse sent from the TTS.

Testing To collect data during sonication, fiber optic hydrophone (Pa Ltd, UK) was used. The fiber optic hydrophone works on the principle of interferometric detection of changes in the optical thickness of a thin polymer film at the tip of the optical fiber sensor downlead. Changes in the thickness may be induced acoustically (through the acoustic pressure) or thermally.

The system is capable of differentiating between the two and making simultaneous measurements of both (Morris et al. 2009). However, in this study, the main purpose was to record the signal during sonication.

Experiment set-up consisted of a water tank filled with degassed water. Gridded surface sensor holder was placed on the top of the water tank (Fig. 67). Fibre-optic hydrophone sensor was mounted on the grid surface. MR Scan was used to find the exact location of the tip of the sensor. This data was used to sonicate to the tip of the fibre-optic sensor by using TTS. Fibre-optic sensor was hard wired to Fibre-optic Hydrophone System control unit. Hydrophone system has its own software to control its hardware. To obtain reading from hydrophone, a computerised scope such as PicoScope (Picotech, UK) was connected via the “AC out” connector on the front panel. The system was designed to have an output impedance of 50 Ohm. PicoScope has two inputs, first is the output of the hydrophone, and second is the monitoring trigger pulse coming from TTS. In order to initiate recording of data, detected by hydrophone, a TTL pulse generated by TTS, was utilised. The time information for the TTL pulse, the ‘execute’ command and ‘stop’ commands were recorded as a tag line in the software (Fig. 68). Deviation during sonication was calculated as t2- t1 (1). Delay after sonication stop button pressed was calculated as shown below (2) length of the signal on PicoScope (t3) minus delay in the start time ( t1) during sonication duration.
$$ \mathrm{Deviation}=\mathrm{t}2-\mathrm{t}1 $$
(1)
$$ \mathrm{Delay}=\mathrm{t}3-\left(-\mathrm{t}1+\mathrm{t}2\right) $$
(2)

Picoscope was programmed using LabView (National Instruments, UK) with 100 ms of time resolution to record the signal simultaneously with the TTL pulse.

Results

With sonication power of 30W for 30 seconds, the deviation of the actual sonication duration from the planned duration was found smaller than 1 second. The time until the transducer stops sonicating, after the stop button was pressed, was calculated as less than 200 ms.

Conclusions

In this study, feasibility of measuring the deviation of the actual sonication and delay after stop button release was tested. The methodology described in this study proves that it is possible to quantify these parameters. With the established methodology, the next step is to quantify the repeatability and reliability of the system with different sonication timings and power values using the Transfusimo Treatment Software (TTS) and the conformal bone system transducer. Transfusimo Treatment System (TTS) is planned to be tested on animals based on these results.

References

[1] Schwenke et al. An integrated model-based software for FUS in moving abdominal organs, Int J Hyperthermia. Vol 31(3): 240–250, 2015.

[2] Morris et aI. A Fabry-Perot fibre-optic ultrasonic hydrophone for the simultaneous measurement of temperature and acoustic pressure”, J. Acous. Soc. Am. Vol 125(6) pp. 3611–3622, June 2009.
Fig. 66 (abstract O53).

Sonication execution with subsonication bank algorithm using Trans-fusimo Treatment Software

Fig. 67 (abstract O53).

Schematic view of experimental set-up

Fig. 68 (abstract O53).

Time diagram for the execution of the sonication with Trans-fusimo Treatment System

O54 Cell transfection, yeast and bacteria transformation with a confocal ultrasound device

Jean-Louis Mestas1, Kamel Chettab4, Gustavo Stadthagen Gomez2, Charles Dumontet3, Bettina Werle2, Cyril Lafon1

1U1032, INSERM, Lyon, Rhone-Alpes, France; 2Bioaster, Lyon, Rhone-Alpes, France; 3U1052, INSERM, Lyon, Rhone-Alpes, France; 4Université de lyon, Lyon, Rhone-Alpes, France

Objectives

Acoustic cavitation can be used for in vitro and in vivo gene delivery as an alternative to viral-based transfection methods.

Gene delivery into expression hosts is one of the first critical steps in recombinant DNA applications. Commonly used methods for DNA delivery into useful biotechnology organisms include, chemically mediated transformation, electroporation and viral transduction. These methods are often cumbersome, long and limited to a single or few specific hosts. This work aims at presenting our development on a confocal ultrasound device for transfecting eukaryote cells with increased efficiency and acceptable viability in a reproducible manner or for delivering functional DNA into Kluyveromyces lactis and Escherichia coli, respectively a yeast and a bacterium.

Methods

This device is based on two piezo ceramic spherical shells placed in a confocal manner (1.1 MHz; Diameter and curvature radius 50 mm, angular gap 90°). This particular configuration is favorable for the initiation and control of cavitation activity. The crossing of the two beams forms an interference patterns that traps the bubbles in the focal zone. The device also integrates a regulation process to control the cavitation activity by adjusting in real time the amplitude of the ultrasound signal as a function of the recorded acoustic response of the cavitation bubbles. With this control loop, the measured activity is within 5% of the desired value. The sonicated volume is placed in 2 ml Eppendorf tubes for cells and yeast (650μl) or in 0.2ml tubes for E-Coli (200μl).

Results

For transfection, the device was evaluated on 11 adherent cell lines and 10 non-adherent cell lines. The presented results are limited on Jurkat and K562 cell lines considered difficult to transfect. The peGFP-C1 transfection efficiency and cell viability were evaluated 24h post sonication. Results show a proportional relation between transfection efficiency and cavitation activities for both cell lines. Optimal transfection rates were 77% and 49% for Jurkat and K562 respectively. The corresponding viabilities were 42% and 84%. These results are comparable to nucleofection method. On a third adherent cell line, A549, this exposure condition gave 80% transfection efficiency for 75% of cell viability.

For transformation, the efficiency was evaluated versus the cavitation index characterizing the cavitation activity level.

Conclusions

A user-friendly and cost-effective ultrasound device was developed. It is well adapted for routine in vitro high-yield transfection and transformation experiments as it does not require the use of any transfection reagent or gas micro-bubbles. It provides a well-adapted method for low cost routine pDNA in vitro delivery for both adherent and non-adherent cell lines yeast and bacteria. This method allows reducing cost for transformation by sonicating bacteria straight in their culture medium. Our results confirm ultrasound as an alternative of non-viral technology for the efficient transient transfection of a wide range of different cells including non-adherent cells or fresh human cells, and the preparation of stably transfected cells.
Fig. 69 (abstract O54).

Stable GFP expression for K562 Cells and e-coli bacteria colonies upon sonication

Fig. 70 (abstract O54).

Experimental set-up

Fig. 71 (abstract O54).

Comparison of sonoporation and electroporation transfection on 2 cell Lines

Fig. 72 (abstract O54).

Yeast and E. coli transformation efficiency evaluation versus sonication parameters

O55 Non-invasive cardiac pacing using image-guided focused ultrasound ex vivo and in-vivo in pigs

Fabrice Marquet1,2, Pierre Bour1,2, Fanny Vaillant1,2, Sana Amraoui1,3, Rémi Dubois1,2, Philippe Ritter1,3, Michel Haïssaguerre1,3, Mélèze Hocini1,3, Olivier Bernus1,2, Bruno Quesson1,2

1IHU Institut de Rythmologie et de Modélisation Cardiaque, Bordeaux, France; 2INSERM U1045 CRCTB, Université de Bordeaux, Bordeaux, France; 3CHU de Bordeaux, Bordeaux, France

Objectives

Currently, no non-invasive cardiac pacing device acceptable for prolonged use in conscious patients exists. The main approach is invasive, employing intravascular catheters, which has associated risks. HIFU can be used to perform remote pacing using reversibility of electromechanical coupling of cardiomyocytes. This technique might be useful in the short term in the clinical settings in various conditions: temporary pacing for bradycardia or any clinical condition with risks of asystole; terminating or examining the inducibility of tachyarrhythmia; screening and optimization of cardiac resynchronization therapy. Here we described an extracorporeal cardiac stimulation device and study its efficiency and safety. We conducted experiments ex vivo and in vivo in a large animal model (pig) to evaluate clinical potential of such a technique.

Methods

Experiments were performed with an MR-guided HIFU platform combining a 1.5T MRI (Siemens Avanto, Germany) and a focused ultrasound device (Image Guided Therapy, France, 256 elements, 13/13 cm aperture/focal, operating at 1 MHz). MR images were recorded using a balanced steady-state free precession sequence (TE/TR/FA/BW = 1.36ms/493ms/80°/1149 Hz.pixel−1, spatial resolution 1x1 mm2, slice thickness 3 mm, 256x256, 40 slices, 3 stacks acquired in transverse, sagittal and coronal orientations) to select the location of the stimulation site and to adjust beam focusing characteristics (mechanical positioning and electronic beam steering). Ex vivo acoustic stimulation threshold was determined performing 756 sonications in the right atrium (83 sonications), the left (431 sonications) and the right ventricles (242 sonications) in 10 ex vivo beating hearts from pigs. In vivo non-invasive stimulation proof of concept was shown performing 314 sonications in 4 anesthetized pigs including 42 sonications without ultrasound contrast agent in the first two animals. The last two animals were injected with ultrasound contrast agents using SonoVue (Bracco, Italy, mean terminal half-life: 12 min, range from 2 min to 33 min). Two consecutive 0.1 mL.kg−1 boli intravenous injection were performed in each animal. Local cardiac electrograms (bipolar measurements) were continuously recorded by three MR-compatible pacemaker leads (CapSureFix MRI Model 5086, Medtronic, MN, USA) inserted into the right ventricle, the left ventricle and the right atrium and connected to a clinical electrophysiology recording system (Bard Inc., NJ, USA). At the end of each in vivo experiment, a navigated delayed inversion-recovery 3D Flash sequence was performed (TE/TR/TI/FA/BW = 3.93ms/714ms/320ms/13°/130Hz.pixel−1, spatial resolution 0.5x0.5 mm2, slice thickness 2.5 mm, 576x576, 52 slices). The animals were injected with 0.2 mmol.kg-1 gadoterate meglumine (Gd-DOTA, Dotarem®, Guerbet, Roissy, France) and scanned 15 minutes post injection. Gross examination of each heart was performed after the heart excision. Histological analysis was performed to assess acute damages screening from acoustic stimulation. Tissue samples of stimulated heart (N=40) as well as control regions (N=24) were collected in 4 ex vivo and 4 in vivo hearts.

Results

Using HIFU it was possible to perform ventricular continuous pacing (A) or to induce ventricular tachycardia (B). Consecutive stimulations of different heart chambers with a single ultrasonic probe was shown, allowing to modify the resulting atrio-ventricular delay (C-D). The results of the 756 stimulation sites performed in the right atrium (RA, 83 sonications), and the left and right ventricles (431 and 242 sonications respectively) in 10 ex vivo beating hearts from pigs were processed to determine stimulation threshold. For each HIFU pulse duration tested ranging from 30 μs to 10 ms, the success of stimulation increases with the acoustic pressure at focus. Two different pressure thresholds were highlighted: one around 4MPa peak negative for HIFU pulse durations above 1 ms and one around 6 MPa peak negative for HIFU pulses ranging from 50 μs to 1 ms (E). The same setup was used in vivo in 4 pigs to show clinical potential (F). Electrophysiological changes were confirmed by arterial pressure modifications (G). The minimal stimulation threshold of 4 MPa negative pressure at the focus (as determined from ex vivo experiments) could not be reached with our current in vivo setup. The maximal peak negative pressure was estimated to be around 2 MPa in situ, due to the limited acoustic window. At this pressure level, stimulation of the LV was observed but with an insufficient success rate. To overcome this limitation and demonstrate in vivo feasibility, ultrasound contrast agents were injected intravenously to enhance HIFU mechanical effects on tissue, hence decreasing the stimulation threshold. Using this protocol, consistent cardiac stimulation was achievable for up to 1 hour sessions in 4 different animals. No damage was observed in inversion-recovery MR sequences performed in vivo in the 4 animals. No signal increase could be seen in the myocardium in the delayed-enhancement MR images that would indicate irreversible injury. Gross pathology and Masson’s staining revealed no differences between stimulated and control regions, for all the ex vivo and in vivo cases.

Conclusions

To the best of our knowledge, this study is the first ex vivo and in vivo proof of feasibility of controlled noninvasive ultrasound-based cardiac stimulation in large animals. The ex vivo characterization demonstrated the potential of this technique in an environment where acoustic parameters were well-controlled and quantitatively determined the stimulation threshold as a function of ultrasound pulse duration and amplitude. The in vivo proof of feasibility performed in large animals showed that this novel technology offers good prospects for clinical developments. Encouraging safety results show that acute stimulation during hour-long sessions did not cause any detectable thermal and mechanical damage under the experimental parameters used.
Fig. 73 (abstract O55).

a Electrophysiological readings of continuous ultrasonic pacing of the heart at 120 min-1 (sinus rhythm: 100 min-1). b Electrophysiological readings of ultrasound-induced non-sustained ventricular tachycardia (165 min-1, sinus rhythm 85 min-1) performed by synchronizing the acoustic emission with the relative refractory period. c-d Example of atrioventricular stimulation with a single ultrasonic probe. Phased array transducer enables consecutive stimulations of the RA (yellow pulse) and the RV (red pulse) with a chosen delay. e HIFU pressure thresholds at the target (peak positive - blue curve- and peak negative -red curve)) vs ultrasound pulse duration to induce ventricle stimulation. f Transverse MR images of the anesthetized pig used during the in vivo proof of concept. g Example of basic electrophysiological and arterial pressure readings. Arterial pressure is reported to prove induction of premature ventricular contraction and non-sustained ventricular tachycardia

O56 A study of the dominant mechanisms of extracorporeal acute cardiac pacing by high intensity focused ultrasound

Amit Livneh, Eitan Kimmel, Dan Adam

Department of Biomedical Engineering, Technion-Israel Institute of Technology, Haifa, Israel

Objectives

Extracorporeal acute cardiac pacing by high intensity focused ultrasound (HIFU) could be a disruptive technology in the field of cardiology. Two clinical applications in which acute cardiac pacing by HIFU may be valuable are: (1) preoperative patient screening in cardiac resynchronization therapy surgery where currently 20-40% of operations fail; (2) Emergency life support, which may prevent an event of cardiac arrest from causing sudden death. Both applications may better morbidity and mortality rates in heart failure patients. While ultrasonic cardiac stimulation was first applied 87 years ago, the mechanisms of ultrasonic cardiac pacing are yet unknown. Our work aims to unveil the dominant mechanisms of HIFU cardiac pacing, using a combined experimental and modeling approach.

Recently, we published results demonstrating HIFU extra systole induction in whole anesthetized rats. Sequences of multi harmony HIFU paced extrasystoles were obtained owing to adequate spatiotemporal control, which employed online ultrasound guidance and real-time vital signs signal processing. An illustration of a sequence of HIFU paced premature ventricular contractions (PVCs) is presented in panel A of the figure below. Visual inspection post pacing showed no indication of gross damage or petechia, histological evaluation didn’t show staining or signs of inflammation 24 hours post pacing. Panel B of the figure below shows a heart post HIFU pacing and histological staining results 24 hours post pacing. Extrasystole induction was demonstrated temporally throughout the entire cardiac cycle beyond the absolute refectory period and spatially across the entire left ventricle. Passive Cavitation Detection (PCD) was applied in conjunction with US imaging on a gel phantom, and on rats. The gel phantom was sonicated with a HIFU pacing sequence, PCD positive cavitation indication was correlated with observed hyperechoic imaging. Similar PCD indications were recorded during in-vivo HIFU pacing, while hyperechoic imaging was not observed. Analysis of these experimental results suggests membrane currents as the dominant cellular level mechanism and cavitation as the dominant ultrasound tissue interaction mechanism. The hypothesis we test here through modeling and simulation is that HIFU induced intramembrane cavitation could induce Premature Action Potentials (PAPs) in a model of a cardiomyocyte by altering the membrane capacitance.

Methods

The Livshitz & Rudy guinea pig LV cardiomyocyte model and O’Hara et al. human LV cardiomyocyte model were adapted to include variable capacitance induced ionic currents and membrane voltage alterations. Numerical simulation in Matlab was applied to calculate the temporal membrane capacitance changes due to simulated HIFU insonation, and the resulting ion and membrane voltage dynamics. The simulated HIFU insonation reconstructed the minimal peak negative pressure that was observed to be required for HIFU pacing in rats.

Results

Numerical simulation results demonstrated HIFU PAP induction throughout the entire diastole (evaluated by the temporal offset from the preceding AP peak of the membrane voltage trace). An illustration is shown in panels C-E of the figure below. The membrane potential is shown in blue. A baseline sinus rhythm was produced by electrical stimulation at 2Hz, the electrical stimulation is noted by the down facing black bars at 0 and 500ms. Ultrasonic pacing was applied at different times during the diastole, the ultrasonic pacing is noted by the upward facing red bars.

Membrane depolarization was gradual, and the ion dynamics composition was similar to that of normal sinus rhythm. The temporal offset between insonation onset and the resulting PAP replicated the in-vivo observations. Moreover, PAP induction was demonstrated to occur also during insonation.

Conclusions

The simulation results of a small animal model reproduced our in-vivo observations. This supports our hypothesis of the suggested dominant mechanisms. The simulation results of a human cardiomyocyte model share similar characteristics and attributes to those of the small animal model, offering the prediction that HIFU pacing could be performed in humans with the same pacing patterns that were applied on whole anesthetized rats.

The presented results offer new insights to the study of HIFU pacing and predict that HIFU pacing may be performed in human subjects without membrane disruption.
Fig. 74 (abstract O56).

Results Highlights. a a Short axis ultrasound (US) imaging of noninvasive HIFU pacing in a rat. The HIFU focus is marked by the curser which is placed on the most dorsal part of the rat’s left ventricle. A sequence of premature ventricular contractions (PVCs), were successively paced by US guided multi harmony HIFU insonation. The PVC sequence is marked on the ECG trace. b A rat heart following HIFU pacing showing no sign of gross damage and an example of Eosin (H&E) staining showing no sign of inflammation 24 hours post pacing. c-e Numerical simulation results of multi harmony HIFU pacing in a small animal model cardiomyocyte demonstrating Premature Action Potentials (PAPs) induction throughout the entire diastole and showing PAP peak voltage was assumed within the same temporal delay from insonation onset as was demonstrated by extracorporeal HIFU pacing in rats. Contractions (PVCs), were successively paced by US guided multi harmony HIFU insonation. The PVC sequence is marked on the ECG trace. b A rat heart following HIFU pacing showing no sign of gross damage and an example of Eosin (H&E) staining showing no sign of inflammation 24 hours post pacing. c-e Numerical simulation results of multi harmony HIFU pacing in a small animal model cardiomyocyte demonstrating Premature Action Potentials (PAPs) induction throughout the entire diastole and showing PAP peak voltage was assumed within the same temporal delay from insonation onset as was demonstrated by extracorporeal HIFU pacing in rats

O57 3D time reversal cavity for histotripsy over a large volume

Justine Robin1, Bastien Arnal2, Mathias Fink1, Mickael Tanter1, Mathieu Pernot1

1Institut Langevin, ESPCI ParisTech, CNRS UMR 7587, INSERM U979, Université Paris Diderot, Paris, France; 2Institut Langevin, ESPCI ParisTech, CNRS UMR 7587, INSERM U979, Université Paris Diderot, Paris, France

Objectives

Ultrasound pulse therapy such as histotripsy or lithotripsy requires focusing very high pressures to mechanically fragment and liquefy tissues. Large spherical transducers are commonly used to achieve these pressures at the focal spot and mechanical steering is then required to treat large regions. Using 2 dimensional arrays of high power transducers is another possibility, but electronic steering is still highly limited by the number of elements that cannot exceed several hundreds for reasons of cost and complexity. In this study, using both numerical simulations and experiments, we have developed a 3-dimentional time reversal cavity (3D-TRC) to focus high intensity pulses over a large volume only using electronic steering, and keeping the number of elements to a minimum.

Methods

We designed a 3D-TRC by enclosing a 3D-multiple scattering medium (MSM) in a reverberating cavity. We used simulations with the k-wave software (pseudo-spectral calculation method, B. E. Treeby and B. T. Cox) and an experimental realisation of our device to optimise its focussing and steering capacity. In both simulations and reality, the cavity was 15x13x20 cm, with steel walls, and filled with water. MSM was either made of steel rods (diameter 0.8mm), or successive metal grids (wire diameter 0.8mm, size of cell 5 mm). Transducers were placed in the back of the cavity, opposite the aperture.

In the simulations, we defined a 119x204x506 matrix, with 0.5 mm grid steps, representing a water volume, in which we placed the steel cavity. A source was placed in front of the cavity in the centre or on the side of the aperture, and emitted a 2-cycle pulse at frequency 1 MHz. Signal was picked up by transducers in the back of the cavity, and stored. Time reversal focusing (TRF) then allowed us to refocus these signals on the initial source point. We explored different kinds of MSM and different sizes and shapes of transducer elements in the cavity. Particularly, we compared the performances of our cavity with either a 128-element linear transducer of high elevational width or an array of 128 square elements of different sizes.

For cavitation experiments, we chose to use 2 high power linear transducers (128-elements, 1 MHz, Imasonics, Besançon, France), placed orthogonally at the back of the cavity, sonicating the MSM with an angle of 60 degrees. The probes were driven by custom multi-channel electronics (Correlec, France). 40 μs US pulses emitted through the cavity were temporally spread to up to 1 ms, picked up by a HGL 200 hydrophone (Onda, Sunnyvale, CA) and stored. Time reversal focusing (TRF) by compressing these signals in space and time then allowed us to reach the needed high negative pressures. Steering the focal spot over a large volume was achieved by moving the hydrophone. We reemitted the reversed signal at a pulse repetition frequency (PRF) of 260 Hz to form a bubble cloud, which was observed using an ultrasound scanner (Supersonic Imagine, Aix-en-Provence, France).

Results

Simulations showed that an array transducer of square elements 3 λ x 3 λ large, with an MSM made of 2 orthogonal very thin rod forests gave the best overall performances in terms of focus quality and steering. Figure 75 shows the focus quality in the centre of the cavity aperture. We thus tried to reproduce this configuration as well as possible in our experiments, but for practical reasons had to work with linear transducers instead.

With our real device in a water tank, hydrophone measurements confirmed the spatio-temporal focalisation of the signal. Observations in a plane 10 cm away from the cavity showed a 1.2 x 1.2 mm focal spot, with a temporal peak less than 2 cycles long. At full power, the peak pressure obtained at the focus was about 40 MPa (linearly extrapolated value). These observations were consistent over a large area (−3 dB area 10x6 cm). The negative pressure obtained was sufficient to achieve cavitation. It was even possible to generate bubble clouds in various spots at the same time by emitting the stored signals corresponding to several locations successively at a PRF of 313 Hz between the different signals and between 16 and 260 Hz overall. Figure 76 shows the bubble clouds formed by the targeting of 2 simultaneous focal spots. We also succeeded in creating lesions in a slice of ham.

Conclusions

Through simulations and experiments, we designed and optimised a 3D-TRC that allowed us to very locally reach the high negative pressures needed to induce cavitation and create lesions in a simple target.

We are confident that we could further improve our experimental results if we fully exploit our simulation results, and move on to a transducer array.
Fig. 75 (abstract O57).

Focal spot observation in 3 orthogonal planes: local maximum pressure is given in arbitrary unit. Focal spot dimensions: 3x3x10 mm. Cross-sections of the time-reversal cavity and MSM are shown on the right in light green

Fig. 76 (abstract O57).

Bubble clouds formed by simultaneous targetting of 2 different focal spots on various positions

O58 Feasibility of transcutaneous volumetric boiling histotripsy ablation of liver and kidney in a pig model

Tatiana D. Khokhlova4, George R. Schade1, Yak-Nam Wang2, Wayne Kreider2, Julianna Simon2, Frank Starr2, Maria Karzova3, Adam Maxwell1, Michael R. Bailey2, Vera Khokhlova2,3

1Department of Urology, University of Washington, Seattle, Washington, USA; 2Applied Physics Laboratory, University of Washington, Seattle, Washington, USA; 3Physics Department, M.V. Lomonosov Moscow State University, Moscow, Russian Federation; 4Department of Medicine, University of Washington, Seattle, Washington, USA

Objectives

Boiling histotripsy (BH) uses millisecond-long pulses of HIFU shock waves emitted at low duty cycle to induce localized boiling in tissue within each pulse. Further interaction of ultrasound with the vapor cavity for the rest of the pulse results in mechanical fractionation of tissue into subcellular debris. Our group is developing BH as a non-invasive treatment of renal and hepatic masses. The feasibility of producing single BH lesions in vivo in an exposed porcine liver has been demonstrated previously. The goal of the present work was to evaluate the feasibility and safety of transcutaneous, volumetric BH ablation of porcine liver and kidney in acute pig studies.

Methods

Pigs (37–40 kg, n=4) were anesthetized and placed on the surgical table in either lateral (for kidney treatment) or supine (for liver treatment) position. A 1.5 MHz HIFU transducer (12-element sector array of 7.5 cm aperture, F#=1.07) with a central opening (2 cm) to allow for ultrasound treatment guidance was attached to a 3D positioning system and submerged in a degassed water bath coupled to the abdomen (Fig. 77). The HIFU focus position, pre-recorded with the ultrasound imaging system, was aligned with the targeted region at the depth of 2.5-4.5 cm from the skin surface. The pulse-average power output threshold for initiating BH at each location was measured by sonicating the focal point with BH pulses at gradually increasing amplitude until a hyperechoic region was observed at the focus, indicating boiling. Prior to the in vivo experiments, similar measurements of threshold output power were performed in freshly harvested ex vivo porcine liver and kidney for comparison to the transcutaneous in vivo setting. The subsequent in vivo sonications were performed slightly above the threshold (10-15% increase in driving voltage). The following treatment parameters were used: pulse duration 5 or 10 ms, pulse repetition frequency (PRF) 1 or 2 Hz (with the duty factor fixed at 1%). A total of 10–30 pulses were delivered per focal point (this number will be further referred to as BH dose), with the focal points spaced 1–1.5 mm apart in a rectangular grid with 0.5-1.5 cm linear dimensions. The BH treatment was not gated by or synchronized with the respiratory motion. Following BH exposure, higher resolution ultrasound assessment of the treated regions was conducted. Necropsy was then performed and the treated portions of the liver and the kidneys were collected for gross and histologic assessment.

Results

Kidney treatments. Lower poles of 7 kidneys were targeted and n=11 volumetric lesions containing cortex, medulla, and renal sinus were created. The transducer driving voltage required to initiate the subcostal transcutaneous treatment in the kidneys was 30 – 50% higher than that observed in the exposed ex vivo tissue; the partially transcostal exposures (30-40% of the beam obstructed by the ribs) required 120-150% larger driving voltage. The respiration-induced motion of the target did not appreciably interfere with the treatment Post-BH, higher resolution ultrasound images revealed well-defined hypoechoic cavities. At necropsy no gross evidence of collateral damage was appreciated within the beam path and no subjects had gross hematuria. On gross inspection of the kidney, small clots were seen within the collecting system in all treated kidneys with regions of petechial hemorrhage surrounding a centrally located fractionated volume of parenchyma. Histologically, all BH exposures produced completely fractionated cortex sharply demarcated from histologically normal untreated tissue (Fig. 78). In the medulla, blood was noted within the collecting ducts with areas of focally fractionated tissue at higher dose exposures (20–30 pulses per focal spot). Within the wall of the collecting system, focal petechial hemorrhage was visualized only at the higher dose exposures without disruption of the wall.

A treatment acceleration strategy was attempted, in which a smaller number (10 vs 30) of shorter (5 ms vs 10 ms) pulses were delivered per focal spot at higher PRF (2 Hz vs 1 Hz) at a larger driving voltage (15% increase). This strategy reduced the overall treatment time 6-fold (resulting in the lysis rate of 3.8 cc/hour), yet achieved the same degree of tissue fractionation as found with the slower treatment.

Liver treatments. Subcostal BH lesions were successfully produced in two out of three livers where treatment was attempted. The threshold for treatment initiation in terms of driving voltage was larger than in the ex vivo porcine liver by 70-200% and was also larger than in the transcutaneous kidney exposures despite very similar treatment depth and body wall thickness.

Most probably, the higher thresholds arose from the aberrative effects of fat within the HIFU beam, as the central section of the body wall contained a much thicker fat layer compared to that overlying the kidney (1.5 cm vs 0.5 cm). The respiration-induced motion of the target was much more pronounced compared to the case of kidney treatments, and led to a noticeable spread of the lesion relative to the planned shape. The hepatocytes in the central region of the lesion were completely fractionated, while at the lesion periphery the treatment effect was less demarcated. Connective tissue structure of the liver lobules, as well as the liver capsule remained intact, consistent with our ex vivo findings (Fig. 78). In the cases where higher power outputs were used (150-200% increase compared to the exposed ex vivo liver), bruising and thermal damage confined to the fatty layer of the body wall were observed.

Conclusions

These data indicate that transcutaneous and partially transcostal volumetric BH treatment of the kidney and liver is feasible in the porcine model. In the kidney, delivering shorter pulses at higher PRF and higher amplitude with constant duty cycle allowed for more rapid, yet equally efficacious tissue fractionation. In the liver, the lesions were successfully generated through a thicker fat layer, without control for respiratory motion. The treatment precision and efficacy can be further enhanced by implementing strategies for phase correction and gating based on respiratory motion. This work was supported by NIH R01 EB7643, K01 EB 015745, NSBRI through NASA NCC 9–58, and Urology Care Foundation.
Fig. 77 (abstract O58).

12-element 1.5 MHz HIFU sector array transducer integrated with an ultrasound imaging probe (ATL P7-4)

Fig. 78 (abstract O58).

Representative photos (left) and tissue sections stained with hematoxylin and eosin (middle) and Masson’s trichrome (right) of porcine liver (top) and kidney (bottom) tissue treated transcutaneously in vivo. Photos show bisected BH lesions (white arrows) with the contents washed out. Histological evaluation reveals regions of completely homogenized tissue (H). Within the lesion, intact collagenous structures (yellow arrow) were present in both organs treated. In the kidney, distinct borders with normal tissue (N) were observed. Small haemorrhagic regions (b) were present in both tissues within and around the border of the lesions.

O59 Non-invasive, rapid ablation of large tissue volume using histotripsy

Jonathan E. Lundt, Steven P. Allen, Jonathan R. Sukovich, Timothy Hall, Zhen Xu

Biomedical Engineering, University of Michigan, Ann Arbor, Michigan, USA

Objectives

Current tumour ablation techniques are typically thermal-based, including radiofrequency (RF), microwave, and high intensity focused ultrasound (HIFU). RF and microwave ablation methods are limited to treating tumours no greater than 3 cm in diameter and at a rate of approximately 2 cm3/minute. While HIFU is capable of treating larger volumes, the treatment duration is excessive. Perfusion-mediated convection (commonly referred to as the “heat sink effect”) presents a major challenge for thermal ablation in highly vascularized tissues. The heat sink effect has been shown to prolong treatment times and result in heterogeneous tissue necrosis. Histotripsy is a noninvasive, non-thermal, ultrasound ablation method that uses high-amplitude, very low-duty cycle focused ultrasound pulses to generate controlled cavitation and thereby mechanically homogenize target tissues into liquid-appearing acellular debris. Our previous in vivo studies have shown that histotripsy is not affected by the heat sink effect and can produce homogenous tissue disruption in the highly vascular liver and kidneys noninvasively through the ribcage and other overlying tissues. Because histotripsy uses microsecond-duration pulses separated by up to seconds of off-time for a given focus, it is possible to electronically steer the focus of a phased array transducer to excite cavitation events throughout a large volume consisting of many overlapping foci during the off-time period. We hypothesize that histotripsy combined with electronic focal steering can achieve rapid ablation of a large target volume. As such, histotripsy can be used to treat tumours that cannot be treated by RF and microwave ablation at a rate exceeding that of these methods. This study presents the first investigation of this hypothesis.

Methods

Histotripsy was applied using a 250 kHz, 256-element phased array transducer with a 30 cm diameter aperture and 15 cm focal distance, generating 1.5-cycle, 6-microsecond acoustic pulses. First, to establish treatment parameters including pulse repetition frequency (PRF) and the number of pulses to deliver, a single-focus lesion was generated in tissue-mimicking phantoms. Tissue-mimicking agarose hydrogel phantoms containing a layer of red blood cells (RBC) allow direct visualization of cavitation and cavitation-induced damage. Cavitation activity and lesion progression during histotripsy treatment were monitored by high-speed optical imaging (Phantom V210, Vision Research) as a function of PRF and the number of pulses applied. Based on the RBC phantom results, ex vivo bovine hepatic tissue was treated by electronically scanning the therapy focus at 200 Hz over 1000 sites (or .2 Hz per focal site). 120 pulses were delivered per site to cover approximately 43 cm3 and 58 cm3 volumes of target tissue (equivalent to spheres 4.3 cm and 4.8 cm in diameter, respectively) over the course of a total treatment time of 10 minutes. The in situ peak rarefactional pressure amplitude was estimated to be 71 MPa at the geometric focus and 49 MPa at the most distal electronic steering foci. Lesion size and morphology were assessed by gross sectioning and magnetic resonance imaging (MRI). Tissue damage was examined by histology using haematoxylin and eosin (H&E) staining of 5-micron sections.

Results

RBC phantom results established that fractionation efficacy degraded at PRFs above 0.2 Hz and that 120 pulses were sufficient to homogenize material within the perimeter of a single lesion. Therefore, 0.2 Hz PRF and 120 pulses per single focus were selected for subsequent ex vivo tissue experiments using histotripsy with electronic focal steering. Morphology of the single lesion was approximated as an ellipsoid with an 8 mm major diameter and a 4 mm minor diameter. Based on the single lesion size, the center-to-center spacing between adjacent steering foci for electronic focal steering treatment was selected to be 2.5/3.2 mm in the lateral plane and 4.1 mm in the axial direction. A total of nine ex vivo bovine liver tissue samples were treated by histotripsy with electrical focal steering. Results from ex vivo experiments show that a completely homogenized and well-defined lesion was generated by histotripsy with electrical focal steering within 10 minutes for all nine treatments. Gross morphology sectioning (Fig. 79a) of ex vivo tissue shows a well-defined region of damage where the treated tissue was liquefied to homogenate. After rinsing away liquid-appearing material, only pale, fibrous structures and vessels larger than about 2 mm in diameter remain. MRI (gradient recalled echo sequence) (Fig. 79b) shows a distinct region of damage with sharp margins of individual foci clearly visible. Histology (Fig. 79c) shows a sharp transition zone (~50 microns) between cells with intact cell walls and a region of scattered cellular material and cell nuclei. For the smaller lateral spacing (2.5 mm), no intact cells remained in the treatment region, while a small fraction of scattered cells were observed in some treatments using larger spacing. MRI 3D volume measurements show the treatment volume to be 43 +/− 6 cm3 for the smaller spacing and 58 +/− 6 cm3 for the larger spacing (mean +/− standard deviation), yielding an average ablation rate of 4.3 and 5.8 cm3/minute, respectively.

Conclusions

Treatment of large and multiple tumour nodules remains a challenge for current tumour interventions, which are mostly thermal-based. This work demonstrates that histotripsy combined with electronic focal steering achieved homogenous and complete ablation of a large target volume at a rate two-fold faster than microwave and RF ablation. Since histotripsy is non-thermal, the treatment should not be affected by the heat sink effect and is expected to remain effective and efficient even in highly vascular organs. With the capability of achieving rapid, homogenous cell disruption, histotripsy has the potential to substantially improve upon current tumour ablation methods.
Fig. 79 (abstract O59).

a Gross morphology of ex vivo bovine liver sample following treatment and irrigation of lesion with tap water. The acoustic axis of the transducer was oriented from left to right with respect to the image. Scale bar: 10 mm (b) MRI of ex vivo sample following treatment. The scan plane is orthogonal to the acoustic axis of the transducer and positioned at the geometric focus. Scale bar: 10 mm (c) H&E histology at 400X total magnification. Intact cells can been seen at the left of the image, fractionated cellular debris and nuclei at the right. Scale bar: 10 microns

O60 Assessment of boiling histotripsy dose in human ex vivo kidneys and renal tumours

George R. Schade2, Yak-Nam Wang3, Tatiana D. Khokhlova1, Philip May2, Daniel W. Lin2, Michael R. Bailey3, Vera Khokhlova4,3

1Medicine, University of Washington, Seattle, Washington, USA; 2Urology, University of Washington, Seattle, Washington, USA; 3Appied Physics Lab, University of Washington, Seattle, Washington, USA; 4Physics, M.V. Lomonosov Moscow State University, Moscow, Russian Federation
Correspondence: Tatiana D. Khokhlova

Objectives

Histotripsy is a pulsed high intensity focused ultrasound (HIFU) technology that mechanically disrupts targeted tissue without a thermal effect. Our group has developed boiling histotripsy approach (BH), a technique that utilizes millisecond-long HIFU bursts to create bubbles at the focus via rapid shock-induced boiling. Interaction of subsequent shocks in the pulse with the ensuing vapor cavity mechanically homogenizes tissue into sub-cellular micron-sized debris in a process involving acoustic fountaining. As a noninvasive, non-thermal approach, BH may have several advantages over existing clinically available thermal ablative technologies for renal masses. Preliminary data suggests differential sensitivities to the effects of histotripsy of specific locations within the kidney, while the sensitivity of human renal tumours to the effects of BH is unknown. The aim of this study was to evaluate and compare the effect of BH on samples of freshly excised human renal tissues and associated tumours ex vivo.

Methods

Freshly excised human kidneys, benign renal tissue, and renal tumour tissue were obtained via IRB approved institutional rapid autopsy and tissue procurement programs. Tissue was obtained from n=11 patients: n=6 whole benign kidneys, n=5 fragments of benign parenchyma, and n=4 tumours (clear cell renal carcinoma (ccRCC): n=2, papillary RCC: n=1, oncocytoma: n=1). All specimens were acquired within 4 hours from death/nephrectomy. Tissue samples were degassed for over 30 minutes in phosphate buffered saline (PBS) and then embedded in low melting point agarose gel. Agarose embedded tissue was then placed in a holder in a bath of degassed PBS. BH exposures were performed under B-mode ultrasound guidance using a 1-MHz 7-element HIFU transducer (aperture 14.7 cm, F#=0.95) with the following pulsing protocol: pulse duration of 10 ms, pulse repetition frequency of 1 Hz, peak focal pressures of p+=88 MPa, p-=17 MPa, shock amplitude of 98 MPa. Single focal volumes within the tumour sample or the renal cortex, medulla, or collecting system were treated at various doses defined here as the number of pulses irradiated into a single focal spot (3–300 pulses/focus). Treated kidneys and tumour samples were evaluated grossly and/or formalin-fixed for histologic assessment with hematoxylin and eosin staining.

Results

Bh pulses produced hyperechoic bubbles at the focus in all tissue types consistent with rapid boiling induced by each pulse. Treatment within the renal cortex and tumour tissue resulted in the development of progressively hypoechoic cavities apparent between pulses, consistent with historipsy treatment effect of homogenizing tissue, while the feedback was less pronounced in the medulla and collecting system. On inspection, tumour tissue appeared more susceptible to the effects of BH than benign tissue; lesions created in tumour tissue with 10 pulses were similar in size to those created with 30 pulses in the cortex (Fig. 80). Histologically, evidence of BH induced tissue homogenization was observed in tumours at much lower dose threshold of f 5 pulses/focus compared to those in benign tissues: 15–30 pulses/focus in the cortex, 45–60 pulses/focus in the medulla, and 90–120 pulses/focus in the collecting system.

Conclusions

BH mechanical ablation of human ex vivo renal tumours is feasible, yielding anticipated tissue homogenization. The observed increased resistance of benign renal tissue to the effects of BH compared to renal tumours, if confirmed in vivo, may help preserve renal function while providing a margin of safety when developing BH for clinical ablation of renal tumours. This work was supported by NIH R01 EB7643, K01 EB 015745, Urology Care Foundation and National Space Biomedical Research Institute (NSBRI) through NASA NCC 9–58.
Fig. 80 (abstract O60).

BH lesion produced in ex vivo human clear cell renal carcinoma (left) and benign human kidney cortex (right) with two different BH doses (10 and 30 pulses, respectively) are similar in size, demonstrating increased tumour tissue susceptibility to BH-induced damage compared to benign tissue

O61 A 200 kHz-1380 kHz multifrequency focused ultrasound transducer for neurostimulation in rodents: numerical study and transcranial in-vitro calibration

Charlotte Constans2, Thomas Deffieux1, Mickael Tanter1, Jean-Francois Aubry3

1Institut Langevin, INSERM, Paris, France; 2Institut Langevin, Université Paris Diderot, Paris, France; 3Institut Langevin, CNRS, Paris, France

Objectives

In order to study the influence of the frequency on transcranial ultrasonic stimulation (TUS), we propose to take advantage of the harmonics of the transducer to investigate a large frequency span with one single element. We calibrated the transducer in the 200kHz-1380kHz range and performed numerical simulations to predict the pressure field that can be generated by this transducer in a rat brain for the same range of frequencies. The transducer was successfully used for neuromodulation experiments on rodents at 320kHz.

Methods

We calibrated a single element transducer with resonance peaks at 200kHz, 320kHz, 850kHz and 1380kHz (H115, Sonic Concepts, Bothel, USA) with a heterodyne interferometer. The pressure in the focal plane after transmission through pure degassed water was measured for each frequency with a power ranging from 18 to 75 electrical watts. Electrical power was generated by a function generator (Handyscope HS5, Tiepie Engineering Sneek, The Netherlands) connected to a 75 W amplifier (75A250A, Amplifier Research) and input voltage and current applied to the transducer were monitored with the channels of the Handyscope (Handyscope HS5, Tiepie Engineering Sneek, The Netherlands).

The acoustic propagation of focused ultrasound was then simulated in an entire rat head in order to investigate the pressure amplitude and spatial distribution as a function of frequency. The simulations were performed with k-Wave [1], a k-space pseudospectral method-based solver. 3D maps of the skull, brain and tissues were extracted from a rat microcomputed tomography scan. Brain and tissues were assumed to have the same sound-speed and density as water, and the transducer was modelled as a spherical section (63mm radius of curvature and 64mm active diameter) with the properties of ceramic. Rather than keeping the resolution constant when investigating the influence of frequency, we fixed the ratio of wavelength to the spatial step to approximately 12 for all simulations. Absorption was taken into account in the skull (2.7dB/cm/MHz) and in the brain (0.37dB/cm/MHz) with a 1.01 power law of frequency. A 230μs-long pulse was simulated, as was used in vivo with the same transducer [2]. Ultrasound propagate in a cone filled with water before entering the rat head, the geometrical focal point being located about 7mm deep from the surface, inside the brain. The simulations are first performed in pure water and compared to the amplitude measured experimentally: the scaling factor is used as a correction factor in order to estimate the absolute pressure in the rat head.

Results

The values of maximum pressure measured in degassed water and simulations in rat brain are summarized in Table 4. One can observe that the same setup is capable of producing more than 1MPa (respectively 1.4MPa) in pure water (respectively in the rat brain) for all frequencies ranging from 200 kHz to 1380kHz.

The maximum pressure in the coronal (top), sagittal (middle) and axial (bottom) planes at the geometrical focal spot is displayed in Fig. 81 for 200 kHz (left) and 850 kHz (right) in a linear scale. As the propagation axis is along the y direction, top figures show the focal plane (view from above the animal) and middle and bottom ones include the propagation path. Stripes originating from standing waves can be seen for both frequencies (Fig. 81, middle and bottom) but are more confined in the 850 kHz simulations. In the axial plane the -6dB area is confined in a 27.5mmx49mm box at 200kHz and 9mm x 25mm at 850kHz.

Conclusions

Simulations show that the same transducer can be used to produce more than 1.4MPa in a rat brain at 200kHz, 320kHz, 850kHz and 1380kHz, which is higher than the threshold for in vivo TUS in rodents [2]. This work paves the way to exploratory work over a large bandwidth with one single experimental setup.

This work was supported by the Bettencourt Schueller Foundation and the "Agence Nationale de la Recherche" under the program “Future Investments” with the reference ANR-10-EQPX-15.

References

[1] B. Cox et al., k-space propagation models for acoustically heterogeneous media: Application to biomedical photoacoustics, J. Acoust. Soc. Am., 2007

[2] Y. Younan et al., Influence of the pressure field distribution in transcranial ultrasonic neurostimulation, Med. Phys., 2013
Table 4 (abstract O61).

See text for description

 

200 kHz

320 kHz

850 kHz

1380 kHz

Maximum pressure in water (MPa)

1.0

1.1

2.7

3.6

Estimated maximum pressure in rat brain (MPa)

2.2

1.4

2.8

3.9

Fig. 81 (abstract O61).

Peak pressure spatial distribution (MPa) in the coronal (top), sagittal (middle) and axial (bottom) planes, in water (left) and in rat brain (right) at 200kHz

Fig. 82 (abstract O61).

Peak pressure spatial distribution (MPa) in the coronal (top), sagittal (middle) and axial (bottom) planes, in water (left) and in rat brain (right) at 850kHz

O62 In vivo study of enhanced chemotherapy combined with focused ultrasound for pancreatic cancer in animal model

Eun-Joo Park, Yun Deok Ahn, Soo Yeon Kang, Dong-Hyuk Park, Jae Young Lee

Radiology, Seoul National University Hospital, Seoul, Korea (the Republic of)

Objectives

As the effects of focused ultrasound (FUS) in anti-cancer drug delivery are widely studied, there is growing interest in the mechanism of how FUS enhances therapeutic effects of drug. In this study, in vivo experiment for pancreatic cancer in animal model was designed to investigate whether non-thermal effect of FUS more effectively enhances the chemo-treatment.

Methods

A pancreatic xenograft model was established by inoculating human pancreatic cancer cells (CFPAC-1) in BALB/c nude mouse. Animals were randomly assigned to the following six groups: control, gemcitabine (GEM) only, FUS1 only, FUS2 only, GEM with FUS1, and GEM with FUS2. Weekly treatments were performed for three weeks and post-treatment tumour size monitoring was followed for five weeks. For FUS treatment groups, animals were sonicated for 20 sec at 1MHz under the guidance of ultrasound images. In combined treatment of GEM and FUS, GEM was administered in IV immediately after the sonication. At the same total acoustic energy, acoustic power and the duty cycles were set in two FUS conditions. Acoustic power was 7.5 W for FUS1 and 80.5W for FUS2. Duty cycle for FUS1 and FUS2 was 50% and 5%, respectively.

Results

Tumour growth rate of animals treated with FUS only (FUS1 & FUS2) was lower than the rate of control group while it was higher than the GEM only group. Animals treated with combination of FUS and GEM showed reduction of tumour growth after two treatments. In FUS2+GEM groups, tumour size reduced until five weeks after the treatment procedure was completed (Fig. 83).

Conclusions

In comparison to longer burst with relatively low acoustic pressure that might have thermal effects on tissue, short burst at high acoustic pressure more effectively control tumour growth in combination with chemo-agent. This result indicates that mechanical reaction induced by FUS can more effectively enhance chemotherapy for pancreatic cancer.
Fig. 83 (abstract O62).

See text for description

O63 Focused ultrasound interventional oncology: local and regional control of advanced malignant disease is here to stay

Vidal-Jove, J.1,4,5; Perich, E.4; Ruiz, A.1; Jaen, A.2; Eres, N.4 ; Alvarez del Castillo, M.3

1Surgical Oncology, HIFU Unit, Hospital Universitari Mutua Terrassa (HUMT), Barcelona, Spain; 2Research Unit, Hospital Universitari Mutua Terrassa (HUMT), Barcelona, Spain; 3Medical Direction Department, Hospital Universitari Mutua Terrassa (HUMT), Barcelona, Spain; 4Interventional Oncology Unit, Institut Khuab, Barcelona, Spain; 5HIFU Onco & Radiology Department, Clínica Santa Elena, Madrid, Spain

Objectives

Interventional Oncology has been proposed as a group of therapeutic procedures that are useful at obtaining local disease control with minimal invasive techniques. We describe our experience with Focused Ultrasound in different tumor settings as well as with different HIFU devices. The 8 years experience of the HIFU Surgical Oncology Unit of Hospital University Mutua Terrassa (Barcelona, Spain) and the 3 years experience of the Interventional Oncology Unit of Institute Khuab Barcelona treating malignant tumors have recently added a new setting for benign & malignant tumors in Santa Elena Clinic in Madrid (Spain). We compare our experience with the reported data in the literature with conventional treatments. We underline some considerations about the role of tumor ablation in the Western World stage.

Methods

From February 2008 to December 2015 we have treated more than 150 tumor cases. Of those, more than 50 cases of non-resectable pancreatic tumors were treated, and we include the first 45 patients from March 2010 to April 2015 to the analysis. All of them also underwent systemic chemotherapy with a standard combination. Devices employed were the JC system as well as the JC200 of USgHIFU from HAIFU Chongqing, China. We have recently added a Echopulse system from Theraclion, France.

Results

The distribution of the 150 cases treated reflects a majority of pancreatic and liver tumors. We specifically analyze the pancreatic tumors. Clinical responses (ablation obtained) were 82% in all cases. We obtained 11 complete responses (25%) at the end of the combined treatment. Major complications included severe pancreatitis (2), skin burning grade III that required plastic surgery (2), duodenal perforation (1). Median Survival is 18 month (6 mo – 3.4 year) and Overall Percent Survival is 33.5% at 5 year follow up.

Conclusions

Focused Ultrasound is an effective and safe Interventional Oncology ablation of benign & malignant tumors. Compared with reported data, it shows survival advantage in non- resectable stage III and IV pancreatic cancer. Interventional Oncology (tumor ablation) needs to be considered as a novel group of therapies along Medical, Radiation and Surgical Oncology and re-defined at the light of this experience.

O64 Polymeric cups as nanoscale cavitation nuclei for active transport and enhanced delivery of nanomedicines into solid tumours

Rachel Myers2,1, James Kwan1, Christian Coviello1, Cliff Rowe1, Calum Crake2, Sean Finn1, Edward Jackson1, Robert Carlisle2,1, Constantin Coussios2, 1

1OxSonics Ltd, Oxford, United Kingdom; 2Institute of Biomedical Engineering, University of Oxford, Oxford, United Kingdom

Objectives

All modern cancer nanotherapeutics, from antibodies and antibody-drug conjugates to oncolytic viruses, suffer from poor passive accumulation, limited penetration and non-uniform distribution in tumours. The present work seeks to exploit microstreaming mediated by sustained inertial cavitation to enable the active transport of biologics from the bloodstream deep into the tumour mass.

Methods

A new generation of nanoscale cavitation nuclei, known as polymeric cups, have been developed both for intra-tumoural and intravenous administration [1]. The cups have a mean diameter of 480 nm, and partially encapsulate and stabilize a single air nanobubble of typical diameter 200–300 nm. Upon exposure to 0.5 MHz ultrasound at in situ peak rarefactional pressures on the order of 1–2 MPa, the cups exclusively produce sustained broadband acoustic emissions associated with inertial cavitation, and generate sustained microstreaming capable of enhancing the transport of co-administered nanotherapeutics unbound to the cups. The cavitating cups can be mapped in real time using a conventional diagnostic ultrasound array and novel Passive Acoustic Mapping (PAM) algorithms capable of identifying sources of broadband acoustic emissions in real time during ultrasound exposure [2]. The present work investigates the usefulness of sustained cavitation mediated by sub-micron cavitation nuclei in enhancing the delivery of different types of oncolytic viruses and antibodies in vitro and to solid tumours in vivo.

Results

In vitro experiments consisted of a flow-through channel in an agar gel as previously described [3]. Little extravasation from the channel was observed in the absence of cavitation or in the presence of non-inertial cavitation mediated by ultrasound contrast agents. By contrast, the generation of sustained inertial cavitation activity mediated by the polymeric cups enables significant penetration of either small molecules, antibodies or viruses to >200 microns away from the vessel wall.

In vivo experiments were carried out using several cell lines and a variety of animal models, including CT-26 in BalbC mice, and HEPG-2 and SKOV in CD1-nude mice, first to quantify any enhancement in delivery and subsequent impact on survival for viruses and antibodies. Cavitation-enhanced delivery was found to enhance oncolytic virus activity in all cases, as quantified by both fluorescene/luminescence and qPCR, by 1–4 orders of magnitude depending on the type of virus being delivered. In separate experiments, the distribution of antibodies to tumours was found to be similarly enhanced, even though the intratumoural antibody dose could not be quantified accurately. Cavitation-mediated delivery significantly inhibited tumour growth both for viruses and for antibodies, and resulted in much more reproducible therapeutic responses across different subjects.

Conclusions

Cavitation-enhanced delivery using sub-micron cavitation nuclei, or polymeric cups, was found to significantly enhance the extravasation, delivery, intratumoural distribution and therapeutic efficacy of both antibodies and viruses for a given systemic dose. The ability to map cavitation activity in real time also offers significant opportunities for real-time monitoring and optimization of successful delivery. Future work will focus on optimizing combined drug and polymeric cup dosing regimes to maximize therapeutic benefit.

References

[1] Kwan et al., Small, 2015

[2] Coviello et al., JASA 2015

[3] Bazan-Peregrino et al., J. Controlled Release 2012

O65 Rapid short pulse (rasp) sequences improve cavitation dynamics for ultrasound therapy

Antonios Pouliopoulos1, Caiqin Li1, Marc Tinguely2, Meng-Xing Tang1, Valeria Garbin2, James J. Choi1

1Bioengineering, Imperial College London, London, United Kingdom; 2Chemical Engineering, Imperial College London, London, United Kingdom

Objectives

Acoustic cavitation – the volumetric oscillation of a bubble due to an acoustic field – is a mechanical force harnessed in therapeutic ultrasound to treat diseases. It can dissolve clots, deliver drugs into cells, deliver drugs across capillaries (e.g., blood–brain barrier opening), and release drugs from liposomes, but it can also cause haemorrhage, kill cells, elicit an immune response, and damage tissue. Current ultrasound parameters have limited control over acoustic cavitation due to our lack of understanding of how microbubbles behave in therapeutically relevant acoustic fields. We introduce a rapid short pulse (RaSP) sequence that has better control of cavitation (e.g., its distribution) than conventional ultrasound parameters. A subset of this design has been previously shown to produce greater therapeutic benefits (e.g., improved drug distribution) in the context of blood–brain barrier opening (Choi et al., PNAS 2011). In order to demonstrate improved dynamics of our RaSP sequence over conventional parameters, we have performed a multi-dimensional analysis of cavitation (e.g., type, magnitude, duration, distribution) using passive acoustic mapping and a high-speed microscopy.

Methods

Traditional therapeutic parameters are composed of long pulses (10–100 ms), which result in microbubble displacement in the axial direction due to a primary acoustic radiation force, clustering due to secondary acoustic radiation forces, coalescence, rectified diffusion, fragmentation, and a variety of other effects due to acoustic cavitation. Many of these effects are undesired sources of mechanical stress that can cause damage. To avoid these effects, we have significantly reduced the length of our pulses (0.01 ms) while increasing the pulse repetition frequency (10,000 Hz). Our pulse shape and sequence exploits the presence of flow by facilitating microbubble movement between pulses. A 0.5MHz focused ultrasound transducer was used to sonicate (PRP: 146-900kPa, PRF: 0.62-10kHz, PL: 5, 25, 50, and 50,000 cycles) microbubbles (SonoVue) flowing within a 800-μm diameter tube, while a ATL L7-4 linear array was used to capture acoustic emissions generated by the cavitation activity. The type, magnitude, distribution, and duration of cavitation activity was analysed using passive acoustic mapping and spectral analysis while a smaller subset of parameters was analysed using a high-speed microscope (5,000 frames per second).

Results

Cavitation persistence during short-pulse excitation increased by 5-fold at low pressures (<150kPa) when compared to a 100-ms long pulse. High pressures and long pulse lengths produced high magnitude inertial cavitation during the first millisecond, which rapidly decreased in energy due to destruction of cavitation nuclei. Cavitation activity was then biased upstream from the focal point due to new microbubbles entering the focal volume. High-speed microscopy observations revealed rapid displacement, clustering, and coalescence at these parameters. Low pressures and short pulse lengths resulted in a more consistent magnitude and distribution of cavitation activity throughout the sequence (figure). High-speed microscopy observations revealed reduced clustering rates and reduced axial displacements.

Conclusions

In conclusion, low-pressure rapid short pulse sequences improved the uniformity of cavitation within the focal volume when compared to long pulses. This improvement was due to the increased lifetime and mobility of the microbubbles within the focal volume. Our demonstration of improved spatio-temporal control of cavitation may improve the efficacy of a wide range of therapeutic applications such as blood–brain barrier opening, sonoporation, and sonothrombolysis, by enhancing therapeutically relevant cavitation dynamics and eliminating unwanted mechanical stress.
Fig. 84 (abstract O65).

Improved magnitude and distribution of cavitation using rapid short pulse (RaSP) sequences. SonoVue microbubbles flowing through an 800-μm-in-diameter tunnel was exposed to a beam of ultrasound (laterial FWHM: 5 mm). Each image represents the lateral distribution of acoustic cavitation activity over time for different exposure conditions (peak-rarefactional pressures: 146–903 kPa, pulse lengths 5, 25, 50, and 16,666 cycles, pulse repetition frequencies: 1,250 Hz)

O66 Early clinical experience of targeted delivery of lyso-thermosensitive liposomal doxorubicin (LTLD) by focused ultrasound to the liver

Paul C. Lyon, Christophoros Mannaris, Michael Gray, Lisa Folkes, Michael Stratford, Robert Carlisle, Feng Wu, Mark Middleton, Fergus Gleeson, Constantin Coussios

University of Oxford & Oxford University Hospitals NHS Foundation Trust, Oxford, United Kingdom
Correspondence: Paul C. Lyon

Objectives

The TARDOX study (Oxford, UK, NCT02181075) is a Phase I first-in-man proof-of-concept study which aims to demonstrate safety and feasibility of targeted drug delivery using lyso-thermosensitive liposomal systems in combination with mild hyperthermia mediated by focused-ultrasound (FUS) applied non-invasively. The primary endpoint of the study is demonstration of enhanced intratumoural delivery of doxorubicin to liver tumours for the same systemic dose of the drug, when given in liposomal form (ThermoDox®) and released locally by ultrasound-induced mild hyperthermia.

Methods

In March 2015, a patient with hepatocellular carcinoma and several liver lesions was recruited to the study and received a single treatment cycle of 50mg/m2 ThermoDox® and FUS. Following an ultrasound-screening session, a single tumour was selected for FUS-targeting, estimated to be of volume 13.7cm3 using a volumetric approximation based on MRI taken the day prior to treatment.

On the day of treatment, the patient was positioned supine over the water bath of the clinical ultrasound-guided extracorporeal FUS device (Model JC200 Focused Ultrasound Tumor Therapeutic System, Haifu Medical). General anaesthesia was induced using high-frequency jet ventilation to minimise respiratory movements of the liver. The water temperature was 14.5°C and patient normothermia was maintained using a controllable heated blanket (Bair Hugger™). Under portable ultrasound guidance, an 18-gauge co-axial needle was placed percutaneously (above the water level) into the core of the target tumour through a sterile field. The central co-axial needle was instrumented with a clinically approved thermistor, interchanged with a core biopsy device according to a treatment protocol. A second, peripheral co-axial needle was placed near to the tumour margin and was used only for thermometry.

Immediately following the 30-minute intravenous ThermoDox® infusion, the JC200 was used to induce hyperthermia in the target tumour using a transcostal approach under conditions of real time thermometry. Thermometry data was acquired via a data acquisition unit (Agilent HP34970A) connected to a PC running a custom LabView client.

The FUS hyperthermia strategy and parameters were selected based on thermometry obtained from previous ex vivo liver tissue using the same system (presented at 3rd European Symposium on FUS Therapy, London 2015). Alignment of the JC200 to the central thermistor was confirmed by low-power test shots. Once the target tumour was contoured through the intercostal space, FUS coverage was planned automatically using single shot mode with a 2mm dot interval, over 11 slices each of 2mm thickness, giving a prescribed volume of 10.5cm3 . The tumour was treated at 50 Watts, 100% duty cycle, row-by-row and slice-by-slice. Rather than heating slices contiguously, a slice separation of four was used in an attempt to dissipate heat more rapidly across the volume and achieve a more uniform bulk temperature rise.

Core tumour biopsies were taken a) prior to drug infusion, b) following completion of drug infusion, and, c) following FUS, for analysis of intratumoural doxorubicin concentration. Biopsy samples were taken in pre-weighed and watertight eppendorfs, which were reweighed before analysis. Samples were frozen at −80°C until the day of analysis, when they were thawed and homogenized. Daunorubicin was added into each sample pot before extraction as an internal standard. Solvent extraction was used to obtain both anthracyclines and their metabolites. A validated high performance liquid chromatography (HPLC) assay with fluorescence detection was used to establish the intra-tumoural concentration of doxorubicin. The ratio of the area under the curve for doxorubicin : daunorubicin was used to calculate the doxorubicin concentrations per gram of tissue. Chromatograms were obtained using Gemini C6-Phenyl guard and analytical columns with Waters 2695 separations Module and 474 fluorescence detector (Watford, UK) with excitation 480nm and emission 560nm.

Dynamic contrast enhanced (DCE) MRI, perfusion CT and 18F-FDG PET-CT imaging was performed the day prior to treatment and at day 17 and 29 post-treatment. An additional DCE MRI was performed the day following treatment. Response evaluation was performed using principles of RECIST & CHOI and the SUVmax metric for the target lesion.

Results

The patient recovered smoothly from anaesthesia and was discharged the following day following clinical review, MRI and blood tests. During the 30-day follow-up period, blood tests were taken at two and four weeks and no adverse events over grade two were reported (NCI CTCAE V4 toxicity criteria).

On the treatment day, following the drug infusion, one complete treatment cycle (354 units) was delivered, followed by a partial cycle (68 units) to maintain hyperthermia over 30 minutes. Post-drug, the JC200 reported an output of 21.1KJ (50W x 422s), taking 32 minutes in real time. Thermometry analysis showed the bulk intratumoural temperature was maintained >40°C for 14m 13s, and >41°C for 50s. During the FUS treatment period, mean and maximal temperatures were 39.8°C and 41.8°C respectively. The peripheral temperature was maintained >38°C for 2m 42s with mean and maximal temperatures of 37.7°C and 38.8°C respectively.

HPLC analysis for the biopsy taken prior to drug administration demonstrated absence of a doxorubicin peak. Following FUS, HPLC revealed a greater than two-fold increase in intra-tumoural doxorubicin concentration, from 2.56 to 5.32μg/g.

Both intra-procedural ultrasound and day 1 MRI demonstrated absence of any changes suggestive of thermal ablation. Subsequent radiological analysis of the target tumour over the four-week period revealed a decrease in attenuation from 75 to 28 Hounsfield units on CT and a reduction in longest axial dimension from 35mm to 25mm on CT and 34mm to 26mm on MRI. PET analysis demonstrated a drop in SUVmax from 4.5 to 3.4-3.8. No such response was seen in control tumours.

Conclusions

The use of LTLD with extra-corporeal FUS hyperthermia for targeted drug delivery in human liver tumours is feasible and may enhance delivery of doxorubicin and its therapeutic efficacy over LTLD alone. Further cases treating tumours of a variety of histological types, size and anatomical locations are needed to support this hypothesis and collate safety data.

O67 RNA-based gene delivery using clinical lithotripter shock waves

Sandra Nwokeoha, Robert Carlisle, Robin Cleveland

Institute of Biomedical Engineering, University of Oxford, Oxford, United Kingdom

Objectives

Nucleic acid-based therapies provide a powerful approach to the treatment of genetic diseases but are challenged by limited delivery. Thus efficient gene delivery strategies are continually being sought. Here mRNA delivery is investigated as, in contrast to the commonly delivered plasmid DNA, mRNA does not require opening of the nuclear envelope, thereby reducing the level of cell injury necessary for transfection. We present for the first time an investigation on the efficacy of lithotripter shock waves (SW) assisted mRNA delivery, based on optimised post-exposure RNA stability and shock wave parameters, in a tissue mimicking system. In addition, we compare SW induced gene augmentation to gene inhibition through the delivery of siRNA.

Furthermore we aimed to determine the transfectability of dissimilar cell types as a function of acoustic pressure and number of SW, to provide insights into the breadth of applicability of SW in the mediation of various gene therapies.

Methods

To optimise SW treatment for optimal transfection, a baseline set of 24 SW conditions was established (n=9 per condition). Three cell lines varying by disease, organ and organism origin were cultured, suspended in continually degassed water and spatially subsumed by the focal zone of a clinical Storz Modulith SLX-F2 electromagnetic shock wave source. The focal volume was measured to be 7.85 mm x 7.85 mm x 42.40 mm. The acoustic exposure parameters comprised peak positive (8.6 – 37.0 MPa) and negative (4.1 – 7.0 MPa) pressures, 125 –1000 shock waves and 1–2 Hz PRF (up to 4 Hz at lower energies due to the capacitance of the shock wave source). Transfectability was assessed as the proportion of permeabilised cells (as assayed by propidium iodide (PI) inclusion) above the proportion of non-viable cells (as assessed using the MTS assay). The structural and biological stability of eGFP RNA was determined by gel electrophoresis and a cell-free in vitro translation method, respectively. Transfections were measured by fluorometry and conducted using a 2 mL tissue mimicking system in which cells embedded in 1% purified agar gels were compartmentalised from a 2.5 mm radius RNA-incorporated channel.

Results

Transfectability was determined at the tested shock wave conditions and 2-D interpolation used to determine the optimal SW dose for maximal cellular uptake per cell type (murine colorectal carcinoma cells shown in Fig. 85). Permeabilisation of normal human kidney cells showed little correlation to SW parameters while poor cell viability recovery at 24 hours for human breast cancer cells resulted in marked cell damage. No statistically significant difference (p<0.05) was found between stabilities of sham RNA and optimal SW exposed RNA. Delivery of eGFP mRNA as measured by expression was enhanced 52-fold by SWs relative to sham treatment (Fig. 86a). A 2-fold decrease in GFP expression was achieved following SW–mediated eGFP siRNA delivery to human breast cancer cells stably expressing GFP (Fig. 86b).

Conclusions

While SWs did not discriminate between normal cells and the characteristically permeability-enhanced cancer cells, optimal SW treatment was cell type specific. Transfection results suggested that SWs may be a mechanism for achieving gene augmentation, by allowing RNA stability and significantly enhanced target protein expression in the absence of external cavitation.
Fig. 85 (abstract O67).

Transfectability of murine colorectal carcinoma cells as a function of the number of shock waves (0 – 1000) and lithotripter energy level 3 (P+ = 8.6 MPa) to 9 (P+ = 37.0 MPa) at 2 Hz

Fig. 86 (abstract O67).

Enhancement of lithotripter SW induced RNA nucleotide delivery. a GFP mRNA delivery using optimal shock wave parameters resulted in a 6- and 52- fold increase in GFP fluorescence intensity relative to sham at 24 and 48 hours respectively. A statistically significant difference (*) was found at the p< 0.05 level. The flourescence were taken at 48 hours. b representative GFP intensities of GFP siRNA shock wave treated cells (+ siRNA/ +SW) scramble siRNA shock wave treated cells (+ scR/ + SW), sham cells (+ siRNA/ - SW) and controls ( - siRNA/- SW). Below are the flourescence intensities across the tissue phantom ROI; the dashed lines demarcate the RNA- incorporated channel. The black arrow represents the direction of SW propagation.

O68 Focused ultrasound enhancement of drug delivery in a pancreatic cancer mouse model

Yak-Nam Wang1, Tatiana D. Khokhlova1, Tong Li1, Navid Farr1, Samantha D'Andrea1, Frank Starr1, Kayla Gravelle1, Hong Chen1, Ari Partanen2, Donghoon Lee1, Joo Ha Hwang1

1University of Washington, Seattle, Washington, USA; 2Philips, Bethesda, Maryland, USA

Objectives

Pancreas cancer remains one of the deadliest of all types of cancer and the most difficult to treat. It is currently the fourth leading cause of cancer death in the United States, and is anticipated to become the second by 2020. Unlike many other cancers, the survival rate for pancreas cancer has not improved substantially, with the five-year survival rate over the past few decades only increasing from 2 to 6%. Current treatment with conventional chemotherapeutics is ineffective due to the presence of extensive stromal desmoplasia and reduced vascular network impeding delivery of chemotherapeutic agents. Our group has evaluated two effects of focused ultrasound (cavitation and mild hyperthermia) to enhance the penetration of chemotherapeutic drug doxorubicin in a genetically engineered mouse model (KPC mouse) of pancreatic ductal adenocarcinoma. This abstract presents the culmination of this work to date.

Methods

All experimental procedures were approved by the Institutional Animal Care and Use Committee of the University of Washington. A KPC transgenic mouse model was used in all these studies. This model closely recapitulates the genetic mutations, clinical symptoms and histopathology found in human pancreas cancer. The tumours have a differentiated ductal morphology with extensive dense stromal matrix and poorly developed vasculature. Doxorubicin was used as the chemotherapeutic agent as a proxy to gemcitabine (current standard of care), due to the ease of detection and the range of clinically approved available forms.

Cavitation. A preclinical focused ultrasound system (VIFU 2000, Alpinion Medical Systems), was used for treatment planning, to apply pHIFU exposures, and monitor cavitation during treatment. The system used either a 1.1 or 1.5-MHz transducer, both of which had a circular central opening of 38-mm diameter fitted with a focused ring-shaped transducer for PCD and an ultrasound imaging probe (C4-12 phased array, center frequency: 7-MHz, Alpinion Medical Systems) for in-line targeting of the tumour. HIFU focal pressures were applied between 1.6–12.4 MPa and 2.2–17 MPa for the 1.1- and 1.5-MHz transducers, respectively. Passive cavitation detectors (PCD), aligned confocally with the HIFU transducers, were used to record broadband emissions from bubble activity during treatment. The following cavitation metrics were calculated from the acquired PCD signals: cavitation probability, cavitation persistence and broadband noise level. Doxorubicin (Dox) was administered during or post pHIFU treatment. The enhancement of drug uptake in the treated area of the tumour was evaluated by multispectral imaging, fluorescence microscopy and high-pressure liquid chromatography (HPLC). The untreated area of the same tumour was used as an internal control. Control animals were not treated with pHIFU.

Mild Hyperthermia. A clinical Magnetic Resonance-guided High Intensity Focused Ultrasound (MR-HIFU) system (Sonalleve V1, Philips, Vantaa, Finland) with a 256-element phased array transducer (focal length 12 cm, frequency 1.2 MHz) was used to apply the focused ultrasound exposures. Therapy planning and real time temperature monitoring was performed using a clinical magnetic resonance imaging (MRI) system (Achieva 3T, Philips, Best, the Netherlands) and a dedicated small animal MR receive coil. The MR sequence was a 2D echo planar fast field echo (FFE-EPI) pulse sequence (TR = 50 ms, TE = 20 ms, flip angle = 20°, voxel = 0.9 x 0.9 x 4.0 mm3, FOV = 100 x 100 mm2, EPI-factor = 7, parallel imaging (SENSE) factor = 2 (RL), saturation bands = 3, dynamic scan time = 1.8 s). Mild hyperthermia treatments were applied (continuous wave ultrasound, acoustic power 7 W) for 10 – 15 minutes, with a binary feedback control algorithm keeping the temperature between pre-defined threshold temperatures (Tmin = 41°C, Tmax = 42.5 °C).

Non-liposomal doxorubicin (Dox) or doxorubicin in the form of a low temperature sensitive liposome (LTSL-Dox) was administered before treatment. The enhancement of drug uptake in the treated area of the tumour was evaluated by fluorescence microscopy and HPLC. Mice not treated with mild hyperthermia were used as controls.

Results

Cavitation. Above the cavitation threshold, the doxorubicin concentration in the treated regions of the tumour was significantly greater compared to the controls. The normalized doxorubicin concentrations were found to be associated with the cavitation metrics (Fig. 87). The pHIFU exposures associated with high cavitation activity resulted in disruption of the stromal matrix and enhanced the concentration by up to 4.5-fold compared to control animals. The increase in drug concentration was supported by both multi-spectral imaging and fluorescence microscopy.

Mild Hyperthermia. The MR-HIFU system enabled localized upkeep of mild hyperthermia within a tight temperature range (41.2 ± 1.3°C) to a target tissue area 6 mm in diameter. Hyperthermia induced by ultrasound increased the median doxorubicin concentration within tumour tissue (Fig. 88) when applied in combination with the systemic administration of low temperature sensitive liposomal doxorubicin (up to 15-fold) or non-liposomal doxorubicin (up to 2-fold) with no significant differences in cardiac levels of doxorubicin. The increase in drug concentration in LTSL-Dox + hyperthermia and Dox + hyperthermia treated animals was supported by fluorescence microscopy. None of the tumours showed damage caused by the application of hyperthermia.

Conclusions

Focused ultrasound can be used to induce cavitation or mild hyperthermia to significantly increase the concentration of doxorubicin into pancreas tumours in the KPC mouse model. The promising results in these studies demonstrate two separate ultrasound-induced techniques that can be used to overcome the barriers to drug penetration in pancreas tumours. This work was supported by the Focused Ultrasound foundation (grant AM01) and US National Institutes of Health (NIH R01CA154451) from the National Cancer Institute (NCI).
Fig. 87 (abstract O68).

Scatter plot of normalized Dox concentration (the outcome) versus cavitation noise level (a) and cavitation persistence (b). The outcomes tend to increase with both the persistence and the noise level. The data from control group (squares), outcomes from the simultaneous treatment group (circles) as well as from sequential treatment group (triangles) are shown. The result lines from generalized estimating equation (GEE) model of cavitation noise level (a) and cavitation persistence (b) are also plotted for the simultaneous pHIFU treatment and Dox administration (dash line) and the Dox administration after pHIFU treatment (solid line)

Fig. 88 (abstract O68).

Box-and-whisker plot of doxorubicin (Dox) concentration in tumours of KPC mice treated with low temperature sensitive liposomal doxorubicin (LTSL-Dox) and non-liposomal doxorubicin (Dox), with and without the application of MR-HIFU. * denotes significance at the p < 0.05 level

O69 Antitumoural effect of bisphosphonates in breast cancer xenografts and bone metastasis is promoted by low-intensity ultrasound

Sophie Tardoski1, Jacqueline Ngo1, Evelyne Gineyts2, Jean-Pau Roux 2, Philippe Clézardin2, David Melodelima1

1LabTAU - U1032, INSERM, Lyon, France; 2Lyos - U1033, INSERM, Lyon, France

Objectives

Bisphosphonates (BP) like zoledronic acid (ZOL) have demonstrated clinical utility in the treatment of patients with bone metastases. However, ZOL exhibits antitumour effects only at high doses incompatible with a clinical use due to renal toxicity. Bisphosphonates exhibit a high affinity for bone mineral, which reduces their bioavailability for tumour cells. We examined if low intensity ultrasound could enhance the effects of a clinically relevant dose of ZOL in experimental breast cancer and bone metastasis murine models.

Methods

A plane transducer working at a frequency of 2.9 MHz was used. The free field acoustic power was adjusted between 8 and 13 watts applied for 30 minutes in order to produce and maintain mild hyperthermia (43°C). These parameters enhance locally the temperature in mice and produce mechanical stimulation without creating cavitation. In vivo experiments were performed in a bone metastases model and on a subcutaneous tumour xenograft model. Animals were randomly assigned to different groups (vehicle, ZOL, US, ZOL+US). Clinically relevant dose of ZOL was used (100 μg/kg). Osteolytic lesions were detected by radiography. Tumour angiogenesis and tumour cells proliferation were assessed by immunohistochemistry. Unprenylated Rap1A, a surrogate marker of the penetration of ZOL into tumour cells, was observed by Western Blotting. A quantification of remaining bisphosphonate in bone after ultrasonic treatment was performed using a fluorescent bisphosphonate (FAM-RIS).

Results

With the acoustic parameters used, no signs of cavitation were found. Temperature in tumours was 42.0 ± 2.8°C during US treatment. No lesion was observed in surrounding tissues. US alone did not have any effect on bone metastasis and tumour outgrowth. In the bone metastasis model, mice treated with ZOL+US had osteolytic lesions that were 58% smaller than those of ZOL-treated animals (p<0.01). ZOL+US also significantly decreased skeletal tumour burden. In the animal model of primary tumours, ZOL+US treatment reduced by 42% the tumour volume, compared with ZOL-treated animals (p<0.01). In all cases tumour angiogenesis and tumour cell proliferation were reduced. Using a fluorescent bisphosphonate, it was demonstrated that US forced the release of bisphosphonate from the bone surface, enabling a continuous impregnation of the bone marrow. Additionally, US forced the penetration of ZOL within tumours, as demonstrated by the intratumoural accumulation of unprenylated Rap1A.

Conclusions

In conclusion, our results demonstrate the potential of low intensity ultrasound as an effective strategy to force bisphosphonate desorption from bone and its penetration through tumour tissue, enabling bisphosphonate antitumour activity (both in bone and outside bone). Our findings made US a promising modality in oncology to trigger anticancer therapy with bisphosphonates.

O70 Characterization of the diffusion properties of different gadolinium-based MRI contrast agents after ultrasound induced blood–brain barrier permeabilization

Allegra Conti1,2, Rémi Magnin1,3, Matthieu Gerstenmayer1, François Lux4, Olivier Tillement4, Sébastien Mériaux1, Stefania Della Penna2, Gian Luca Romani2, Erik Dumont3, Benoit Larrat1

1CEA/DSV/I2BM/NeuroSpin, Gif sur Yvette, France; 2Department of Neuroscience, G. D'Annunzio University, Chieti, Italy; 3Image Guided Therapy, Pessac, France; 4Université Lyon 1, Lyon, France

Objectives

The in vivo characterization of Gadolinium (Gd) based MRI Contrast Agents (MR-CA) diffusion within brain extracellular space is of great interest for the understanding of drug transport in brain parenchyma in the framework of new pharmaceutical developments for Central Nervous System diseases. We present here a new method to study the diffusion process of different MR-CAs after a transient and local Blood–brain Barrier (BBB) permeabilization induced by ultrasound. By estimating the Free Diffusion Coefficients from in vitro studies, and the Apparent Diffusion Coefficients from in vivo experiments, an evaluation of the tortuosity (λ) in the right striatum of 11 Sprague–Dawley rats has been performed.

Methods

Four Gd-chelates with different hydrodynamic diameters (dH) were tested: three commercially available MR-CA (MultiHance, Gadovist and Dotarem) and a new class of Gd-based nanoparticles (AGuIX, Nano-H). These latest compounds are composed of a core of polysiloxane, grafted with two or three Gadolinium chelates. They are sufficiently small (dH < 5 nm) to escape hepatic clearance. Diffusion Light Scattering (DLS) measurements were performed to estimate the hydrodynamic diameter of all compounds (Table 5). The MRI acquisitions were performed with a 7T/90 mm Pharmascan scanner (Bruker). The contrasting power of each CA is characterized by its longitudinal relaxivity, r1. To evaluate the CAs’ longitudinal relaxivities at 7T and 37°C, bundles of tubes containing different CA-concentrations in 0.3% w/w agar gel were prepared. The T1 values of these tubes were measured using an IR-FGE sequence, by fitting the signals as a function of TI (S(TI) = | A-Bexp(−TI/T1*) |, where T1= T1*x[B/A–1]). Figure 89(a) shows an IR-FGE MR image for one particular TI, and Fig. 89b shows the signal fits in each tube. From the T1-maps (Fig. 89c) the longitudinal relaxivity was extracted by the linear fit: R1 = 1 / T1 = 1 / T10 + r1 x [CA], where T10 is the T1 of the media without Gd. Relaxivity values r1 are summarized in Table 5 for all compounds. The Free Diffusion Coefficients (DFree) of these compounds were then estimated by injecting 10 μL of a 5 mM solution in a tube filled with 0.3% w/w agar gel. The diffusion was followed for 1 hour by acquiring five T1-maps. The tubes were kept at 37°C during the acquisition. A T10-map acquired before the injection was used as a reference.

MR-CA concentration maps were then calculated from T1 maps using the previous equation. On each CA map, the following 2D Gaussian function (3) was fitted:
$$ \mathrm{C}\mathrm{A}\left(\mathrm{x},\mathrm{y}\right)=\mathrm{A}\ast \exp \left(-\mathrm{a}{\left(\mathrm{x}-{\mathrm{x}}_0\right)}^2-\mathrm{b}\left(\mathrm{x}-{\mathrm{x}}_0\right)\left(\mathrm{y}-{\mathrm{y}}_0\right)-\mathrm{c}{\left(\mathrm{y}-{\mathrm{y}}_0\right)}^2\right) $$
(3)

where A is the Gaussian amplitude and (x0,y0) are the coordinates of its center along the absolute axes (x,y). a, b and c are functions of the Gaussian spreads (σx and σy) along its main axes (X and Y) and of the angle θ between these axes and (x,y).

By taking σ2 x and σ2 y as the molecules mean square displacements along X and Y, the diffusion coefficients along these axes, Dx,vitro and Dy,vitro, are given by Dx,y,vitro = σ2 x,y/2t, where t is the time after injection, i.e. the diffusion time. The Free Diffusion Coefficient was then calculated as the mean value of Dx,vitro and Dy,vitro. Hydrodynamic diameter of the CA was deduced from Dfree using the Stokes-Einstein formula (4):
$$ {\mathrm{D}}_{\mathrm{Free}}=\mathrm{k}\mathrm{T}/\left(3{\uppi \upeta \mathrm{d}}_{\mathrm{H}}\right) $$
(4)

where k is the Boltzmann constant, T is the temperature in Kelvinand η is the viscosity of the agar gel (6.92 x 10−4 Pa.s). Focused ultrasound induced BBB permeabilization was then performed in the right striatum of 11 Sprague–Dawley rats (120 g, Janvier, France). To do so, we used our previously developed motorized MR guided transcranial FUS system (Image Guided Therapy, France) [1]. It enables to position the ultrasound beam in the rat brain at high precision. A 1.5MHz, 8 channel MR compatible concave transducer was calibrated and mounted on the system for this study (Imasonic, France). It was coupled to the shaved rat head with echographic gel. After reference anatomy, T1 weighted and T1 mapping scans, the BBB opening protocol was performed as follows: sonovue microbubble intravenous injection (Bracco, Italy), pulsed ultrasound were shot at 0.5MPa (3ms/100ms for 60s). Then, 200μL of 5mM MR-CA were injected intravenously. The Gd chelates diffusion starting from the BBB disruption site was followed by repeatedly acquiring T1-maps for about 1 hour, as for the in vitro measurements. Animals were kept under general anesthesia during the whole procedure (1.5% isoflurane).

The same Gaussian fitting procedure was applied and the Apparent Diffusion Coefficients (ADC) of all compounds in the striatum were estimated as the average ADC = (Dx,vivo + Dy,vivo)/2.

Results

Figure 89 presents the steps to measure CA longitudinal relaxivities. Figure 90a shows examples of in vitro CA-maps obtained for MultiHance. The fitted 2D-Gaussian functions are presented in Fig. 90b, whereas in Fig. 90c the linear fit on their spreads is plotted. Figure 91 shows an example of in vivo dataset acquired by injecting Dotarem in one rat after focal BBB disruption. In Fig. 91a the original CA maps acquired within 66 minutes after the CA injection are pictured, and their respective Gaussian fits are shown in Fig. 91b. The linear fit over these Gaussian spreads is given in Fig. 91c. As can be noticed in Table 5, both DFree and ADC are decreasing with increasing hydrodynamic diameters (Dotarem > Gadovist > MultiHance > AGuIX). Furthermore, quantitative values of hydrodynamic diameters deduced from Dfree measurements are really consistent with DLS measurements. The ADC values have been used to estimate tissue tortuosity λ=(DFree/ADC)0.5 (Table 5), showing a very good agreement with the tortuosities evaluated with more standard techniques [2].

Conclusions

The agreement between the values of λ found after the blood–brain barrier permeabilization and the known values typical of healty brain tissue confirms the validity of this method to estimate the ADC values in the tortuous regime, but also that the diffusion properties of the tissue are not altered by the ultrasound induced BBB permeabilization protocol unlike by direct intracerebral injection [3]. This should be taken into account for CNS drug development since pharmacodynamics might be modified by direct injection.

References

[1] Magnin et al., ISTU conference proceeding 2014

[2] Nicholson et al.,Trends Neurosci. 1998, 21: 207–215

[3] Marty et al., CMMI 2013, 8: 12–9
Table 5 (abstract 70).

Extracted parameters for each compound (Dotarem, Gadovist, MultiHance and AGuIX). Notably, the apparent dH estimated from the Stokes-Einstein equation is in agreement with DLS measurements, and both DFree and ADC values decrease when molecular size increases. The tortuosities are consistent for all compounds and in agreement with literature

 

Number of Rats

r1 at 7T (s−1·mM −1)

DFree (10−11 m2·s−1)

ADC (10−11 m2·s−1)

Stokes Einstein dH (nm)

DLS dH (nm)

λ

Dotarem

3

4.7

45

16

1.5

1.4

1.7

Gadovist

3

5.5

39

15

1.7

1.8

1.6

MultiHance

3

7.3

28

11

2.3

2.3

1.6

AGuIX

2

8.0

11

6

5.8

3.5

1.5

Fig. 89 (abstract O70).

For each compound the r1 value was estimated by fitting the IR-FGE signals as a function of TI(S(TI) = |A ‐ B × exp(‐ TI/T1*)|, where T1 = T1* × [B/A ‐ 1]). Fig. a shows an IR-FGE MR image for one particular TI, and Fig. b shows the signals fits in each tube. From the T1, and Fig. b shows the signal fits in each tube. From the T1-maps (Fig. c) the longitudinal relaxivity was extracted by the linear fit: R1 = 1/T1 = 1/T10 + r1 × [CA]

Fig. 90 (abstract O70).

In vitro diffussion of MultiHance: a concentration maps and their repective 2D Gaussian fits (b) acquired during 1 hour CA injection in 0.3% w/w agar gel. Fig. c shows the squared Gaussian spreads as a function of time and their fits (σ 2 x,y = 2t × Dx,y vitro)

Fig. 91 (abstract O70).

In vitro diffusion of Dotarem in the frontal hemisphere of a rat brain after ultrasound-induced BBB permeabilization (a) and the corresponding 2D Gaussian fits of CA-sports (b). Fig. C shows the linear fits of the squared Gaussian spreads as a function of time after CA injection

O71 Closed-loop control of targeted drug delivery across the blood–brain barrier

Tao Sun1,2, Chanikarn Power1, Yong-Zhi Zhang1, Jonathan Sutton1, Eric Miller2, Nathan McDannold1

1Department of Radiology, Brigham and Women’s Hospital, Harvard Medical School, Boston, Massachusetts, USA; 2Department of Electrical and Computer Engineering, Tufts University, Medford, Massachusetts, USA

Objectives

Microbubble-mediated focused ultrasound (FUS) can induce localized and reversible blood–brain barrier disruption (BBBD), aiding in targeted drug delivery to the brain. Acoustic cavitation is well-accepted as the primary mechanism in opening the BBB. In addition to its therapeutic potential, inducement of inertial cavitation by FUS can result in permanent damage to the brain tissue. Previous studies have shown the utility of cavitation detection for monitoring all bubble activities during treatment; however translation of FUS induced BBBD requires the development of a closed-loop, real-time control system that can tailor the opening while simultaneously keeping the brain damage-free. Here, we propose an acoustic emissions-based controlling paradigm that can not only modulate the BBBD outcomes based on the feedback from stable cavitation responses (harmonic emission, HE), but also suppress the likelihood of brain damage by monitoring the inertial cavitation components (broadband emission, BE).

Methods

This controlling system has been designed and tested in a preclinical dual-transducer setup. Two FUS single-element transducers were driven at different frequencies (F = 274.3 KHz, F = 30 Hz), and oriented at 120° to create a sub-centimeter focal depth-of-field. Cavitation activities were monitored using a passive detector (central frequency: 650 kHz).

Injections of OptisonTM into the tail vein in rats (n = 43) were performed with a computer-controlled injector that constantly rotated the syringe to keep the microbubbles mixed. HE and BE were monitored in real time and used as the basis for control of the next FUS pulse. The impact of multiple pulse parameters, including acoustic pressure, pulse repetition frequency (PRF), and total duration, was studied. To minimize the potential damage caused by inertial cavitation, the exposure level was reduced if BE was detected and terminated in the event it crossed a set threshold. The performance of the control system was assessed by BBBD in rats in vivo. Trypan Blue dye was injected as our model drug for visualizing and quantifying BBBD by fluorescent imaging 1-h post FUS treatment.

Results

A pilot study investigated the microbubble-dose effects on the linearity of the HE-pressure dependence and on the inertial cavitation threshold (ICT). The linearity of the HE was confirmed, and the ICT was found to decrease as microbubble dose (100, 200, and 400 μl/kg) was augmented. Assessment of the controller performance demonstrated that: 1) Control of HE is possible while keeping BE-free; 2) The use of microbubble infusion after an initial bolus allows the proportional controller to rapidly converge and prevents the decline of emissions, thus reducing the likelihood of BE due to the limited number of cavitation nuclei at the later stage of sonication; 3) A higher PRF (1 vs. 4 Hz) significantly enhanced the percentage of received emissions within the preset goal (Good Burst Rate, GBR), both in bolus injection mode (2-fold, P<0.05) and in infusion injection mode (7-fold, P<0.001); 4) Using an infusion and a PRF=4 Hz, the GBR for HE of 85.6% ±4.0% during 120-pulse sonications was achieved. Statistical comparison of GBR between the 60-, 90- and 120-pulse sonication showed no significant difference (P>0.05), suggesting the robustness of the controller. Total HE was exponentially correlated with the radiant efficiency of epi-fluorescence emitted from the Trypan Blue dye (R 2 = 0.78).

Conclusions

We have designed a real-time, closed-loop FUS controller based on passive cavitation detection for microbubble-mediated ultrasound therapy. The controller performance has been optimized through microbubble infusion and changing PRF. Controlled BBBD in rat in vivo was achieved based on the guidance of the present controller.

O72 Two-dimensional array synthesis using raster-scanned single hydrophone in application to acoustic holography and ultrasound imaging

Oleg Sapozhnikov1,2, Sergey Tsysar1, Petr V. Yuldashev1, Vera Khokhlova1,2, Victor Svet3, Wayne Kreider2

1Physics Faculty, Moscow State University, Moscow, Russian Federation; 2Center for Industrial and Medical Ultrasound, Applied Physics Laboratory, University of Washington, Seattle, WA, United States; 3N.N. Andreyev Acoustics Institute, Moscow, Russian Federation

Objectives

Two-dimensional (2D) receiving arrays with ~104 total number of elements would serve as a powerful tool for ultrasound imaging, especially in the situations when the region of interest is imaged through inhomogeneous layers like skull or ribs. While 2D arrays with such a large number of elements are not yet practically available, their imaging capabilities can be studied by replacing them with synthetic 2D receiving arrays that are made using a single hydrophone that is raster scanned by a computer-controlled positioner. The goal of this study is to illustrate such an approach experimentally by using a synthetic 2D array in two applications: acoustic holography of therapeutic ultrasound sources and ultrasound imaging through a skull phantom.

Methods

To create a synthetic array, a miniature ultrasound probe was moved point-to-point using a computer-controlled positioning system. Measurements were made in degassed water using several experimental arrangements, in which the ultrasound probes were either capsule or needle hydrophones with sensitive diameters of 0.15, 0.2, or 0.5 mm. The scans were executed using positioning systems with stepper motors and linear slides that provided a resolution of several microns per step. In each scan, the hydrophone was sequentially located at the nodes of a square grid with a pitch less than half the wavelength. Typical size of the scanning region was 100×100 points, which corresponded to the number of elements of the corresponding virtual synthetic array. At each hydrophone position, an ultrasound source emitted a tone burst which was received by the hydrophone. Then, a hydrophone was moved to a new location and the operation of emission-reception was repeated. The pressure waveform at each location was recorded using a digital oscilloscope and transferred to a computer.

Results

Synthetic arrays were used for two applications: The first was acoustic holography, which was shown in our previous work to be a powerful technique for characterizing ultrasound sources and the fields they radiate [see 1,2]. Beyond the CW version of holography that is appropriate for most regimes of ultrasound therapy, transient holography is directly relevant to imaging applications and therapies like histotripsy [3]. Here, a transient hologram was detected by a synthetic array (150×150 size, 0.7 mm pitch) in order to characterize a piezoelectric HIFU source (10 cm diameter, 10 cm focal length) excited by a 7-cycle tone burst at a frequency of 1 MHz. The reconstructed vibration velocity magnitude is shown in Fig. 92.

The second application of synthetic arrays was ultrasound pulse-echo imaging through an inhomogeneous layer mimicking a human skull. Transcranial ultrasound imaging remains problematic due to severe aberrations caused by the skull. Wide-aperture 2D arrays can help to achieve usable imaging resolution by compensating for aberration effects. In the experiments, a skull phantom was made from epoxy resin mixed with aluminum oxide powder. The phantom had the following parameters: density 1.4 g/cm3, longitudinal velocity 2.6 mm/μs, shear velocity 1.3 mm/μs, and absorption coefficient 4 dB/cm at 1 MHz, 7 dB/cm at 2 MHz. The phantom thickness was made nonuniform with one side being flat and the other side having profile variations similar to human skull. To simulate the “flash-mode” imaging regime, the skull phantom was insonified from the flat side by a short 2-MHz tone burst emitted by a broadband wide-aperture (several cm diameter) flat source. A needle hydrophone was raster scanned in a plane region proximal to the skull phantom. The corresponding synthetic array was of 100×100-element size and 0.5 mm pitch. Several mm-sized scatterers were placed in water at 3–4 cm distance from the other side of the phantom. The imaging consisted of two steps. In the first step, the skull phantom thickness was mapped using echo arrival time differences between the front and back surfaces (Fig. 93). Then the 3D image was built based on a delay-and-sum algorithm. The image was built both without and with account for the presence of the inhomogeneous skull phantom. Typical image improvement can be seen in Fig. 94: the lateral resolution was significantly improved when the aberrations were accounted for in the procedure.

Conclusions

Synthesizing a 2D array with a large number of elements can be effectively done even with a single hydrophone. To mimic the array, a hydrophone placed in the desired array elements’ locations by a computer-controlled positioner can be used. Array synthesis is possible and effective if the acoustic field under study can be generated repeatably with high accuracy. Capturing acoustic field measurements in 2D provides detailed information about 3D fields in both CW and transient regimes. This information has practical utility in both therapeutic and imaging applications. The work was supported by the Russian Science Foundation grant no. 14-15-00665 and NIH R21EB016118.

References

[1] Sapozhnikov et al. Acoust. Phys. 49(3), 354–360 (2003)

[2] J. Acoust. Soc. Am. 138 (3), 1515–1532 (2015)

[3] Sapozhnikov et al. Acoust. Phys. 52(3), 324–330 (2006)
Fig. 92 (abstract O72).

Piezoelectric transducer (left) and distribution of the particle velocity magnitude on the transducer surface while operating in the transient regime (right)

Fig. 93 (abstract O72).

Skull phantom thickness pulse-echo measurements. Left: typical B-mode image in a transversal plane. Right: phantom back-side profile reconstructed from the pulse-echo measurements

Fig. 94 (abstract O72).

B-mode image of a 2-mm diameter spherical scatterer placed behind the skull phantom. Left: image constructed without account for the inhomogeneous layer. Right: image reconstructed with account for the layer

O73 Ultra-high speed imaging and modelling of shock wave interactions with cells

Dongli Li, Antonio Pellegrino, Nik Petrinic, Clive Siviour, Antoine Jerusalem, Robin Cleveland

Department of Engineering Science, University of Oxford, Oxford, United Kingdom

Objectives

Visualise and model cell deformation under the action of shock waves in order to optimise therapeutic efficacy while minimising side-effects during treatment.

Methods

Shock waves from a clinical shock wave source (Minilith, STORZ) were applied to a tissue-mimicking phantom with embedded cells. The pressure profile of the shock wave was measured inside the phantom using a PVDF needle hydrophone and the induced strain rate was estimated to be 105-106 s−1. The deformation of individual cells were tracked using an ultra-high speed camera with frame rate up to 20 Mfps. The cell response was analysed by its area and perimeter change over time. A Finite Element (FE) model was developed to compare to the experimental findings with the same setup using a combined non-linear fluid and hyper-viscoelastic framework.

Results

Under the compression phase of shock waves, cells showed 3-5% of area decrease and 1-2% of perimeter reduction, whereas under the tension phase, cells showed 15-20% of area increase and 6-8% of perimeter rise. Simulation results matched with the experimental findings by proposing a new constitutive model differentiating the compressive behaviour from its tensile counterpart.

Conclusions

The results of this study suggest that at high strain rate the cell appears as much stiffer in compression than in tension because of the intrinsic deformation mechanisms. The accurately characterised cell properties can thus help to predict cell response in order to optimise the therapeutic efficacy of shock wave applications such as lithotripsy, orthotripsy and cancer treatment.
Fig. 95 (abstract O73).

Imaging processing for the ultra-high speed images

Fig. 96 (abstract O73).

Comparison of cell contours during shock wave interactions

O74 Comparison of derating methods for nonlinear ultrasound fields of diagnostic-type transducers

Peter V. Yuldashev1, Maria Karzova1, Bryan W. Cunitz2, Barbrina Dunmire2, Wayne Kreider2, Oleg Sapozhnikov1,2, Michael R. Bailey2, Vera Khokhlova1,2

1Physics Faculty, Moscow State University, Moscow, Russian Federation; 2CIMU, Applied Physics Laboratory, University of Washington, Seattle, WA, United States

Objectives

There are therapeutic and diagnostic uses of imaging probes, which benefit from exceeding the Mechanical Index limits of diagnostic ultrasound and support that these benefits occur without negative bioeffects. Without imbedded software restrictions, the in situ pressure levels of these devices can exceed the typical diagnostic limits on Mechanical Index and spatial peak pulse average intensity (ISPPA). When calibrating imaging probes at these levels in water, nonlinear propagation effects are present, which complicates the derating process for estimating in situ fields. Different derating approaches have been proposed to predict pressures in tissue from measurements in water. One conventional derating method is to scale the focal pressures and another method is to scale the source amplitude to compensate for linear losses on the way to focus. This second method is described as nonlinear derating and has been shown to provide accurate results for strongly focused therapeutic transducers. However, applicability of these derating methods to diagnostic probes operating at therapeutic intensities is still in question. Here, the derating methods were tested for a diagnostic probe used in kidney stone propulsion technology.

Methods

A standard diagnostic ultrasound curved array probe operating at 2.3 MHz (C5-2, Philips Ultrasound, Andover, MA, USA) was considered. The array comprises 128 elements; however, the results presented hereafter were obtained by considering 64 central elements of the array to be active (Fig. 97). In the azimuthal plane the focus can be steered electronically, while a cylindrical acoustic lens focuses the field at a fixed depth in the elevation plane.

A combined measurement and modelling approach was used to establish an equivalent source boundary condition for nonlinear simulations of the array field in water based on the 3D Westervelt equation [1]. Simulations were performed for propagation entirely in water and in the presence of a tissue mimicking phantom placed at a distance of 1 cm in front of the probe surface. The acoustic properties of the phantom were set the same as in water, except for the frequency dependent absorption of α0 = 0.5 dB/cm/MHz with power exponent n = 1.2 and corresponding dispersion that were set according to the properties of the phantom.

Two derating methods were tested to estimate the in situ (z = 50 mm) pressure field in tissue from the waveforms simulated in water. Derated waveforms were then compared with direct simulation results in tissue. In the conventional derating method, the pressure field calculated in water at the focus was multiplied by the absorption exponent accounting for the propagation distance of 40 mm in tissue to the focus. In the nonlinear derating method, the pressure amplitude at the focus in tissue was assumed to be the same as in water for the lower source voltage scaled with the same absorption exponent value.

Results

Axial distributions of the peak positive and peak negative pressures are shown in Fig. 98 for several output voltages in the free field in water (a) and in the presence of the tissue phantom (b). The focal lobe of the probe (20 mm long) is relatively large in comparison with the focal length of 50 mm because the transducer has a relatively low linear focusing gain (9.3). Therefore, at high power outputs nonlinear propagation effects accumulate over the long propagation distance and are not localized near the focus as is the case for strongly focused therapeutic sources.

Peak positive and peak negative pressures at the focus, z = 50 mm, as functions of source voltage are shown in Fig. 99. The nonlinear saturation of the peak positive pressure is clearly seen for propagation in water (black curve) and in tissue (blue curve), though at higher voltage levels. The conventional derating process of scaling focal pressures is illustrated by the green curve in Fig. 99. As denoted by the vertical dashed arrows, this method overestimates peak positive pressure at moderate voltages (up to 50 V) and underestimates it at higher voltages. Nonlinear derating is illustrated by the red curve. As shown by the horizontal arrows, peak positive pressures are significantly overestimated (by up to 50%) for source voltages higher than 20 V. For lower voltages, the nonlinear derating matches results in tissue within 10% of accuracy. Peak negative pressure magnitudes estimated by derating methods are smaller than those obtained in direct numerical simulations. Peak negative pressures predicted by the conventional derating method can be 50% smaller than in simulations, while the discrepancy remains less than 20% for the nonlinear derating method.

Focal waveforms obtained in simulations in tissue (blue curve) and using the derating methods (red and green curves) are compared in Fig. 100 for 55 V (a) and 90 V (b) outputs. At 55 V the nonlinear derating method predicts 40% higher peak positive pressure than simulations, while peak negative pressures are in closer agreement. The waveform resulting from conventional derating is fortuitously close to the simulated waveform. At 90 V all waveforms are significantly different; the peak positive pressure obtained in simulations is approximately in the middle between the results of the two derating processes.

Conclusions

Nonlinear acoustic fields generated by a diagnostic ultrasound probe are simulated in water and in a tissue phantom using the 3D Westervelt equation. Two derating approaches were applied to estimate the pressure field in tissue using the results obtained in water. It was shown that the conventional derating method can either overestimate (up to 50%) or underestimate (up to 25%) peak positive pressure depending on the source voltage, while it underestimates peak negative pressures by up to 50%. The nonlinear derating method provides accurate results at low intensities (here up to 20 V); however, it overestimates peak positive pressures by up to 50% at higher intensity levels. These simple derating procedures therefore cannot substitute direct numerical modelling to provide reasonable accuracy for nonlinear in situ pressures for diagnostic probes. The study was supported by the grants RSF 14-12-00974, NIH EB007643 and DK043881.

Reference

[1] Karzova et al., AIP Conf. Proc. (1685) 040002–1, 2015
Fig. 97 (abstract O74).

Geometry of focused ultrasound beam produced by C5-2 array probe with 64 active elements

Fig. 98 (abstract O74).

Axial distributions of the peak positive and peak negative pressures for a