The impact of vaporized nanoemulsions on ultrasound-mediated ablation
© Zhang et al.; licensee BioMed Central Ltd. 2013
Received: 9 August 2012
Accepted: 13 January 2013
Published: 25 April 2013
The clinical feasibility of using high-intensity focused ultrasound (HIFU) for ablation of solid tumors is limited by the high acoustic pressures and long treatment times required. The presence of microbubbles during sonication can increase the absorption of acoustic energy and accelerate heating. However, formation of microbubbles within the tumor tissue remains a challenge. Phase-shift nanoemulsions (PSNE) have been developed as a means for producing microbubbles within tumors. PSNE are emulsions of submicron-sized, lipid-coated, and liquid perfluorocarbon droplets that can be vaporized into microbubbles using short (<1 ms), high-amplitude (>5 MPa) acoustic pulses. In this study, the impact of vaporized phase-shift nanoemulsions on the time and acoustic power required for HIFU-mediated thermal lesion formation was investigated in vitro.
PSNE containing dodecafluoropentane were produced with narrow size distributions and mean diameters below 200 nm using a combination of sonication and extrusion. PSNE was dispersed in albumin-containing polyacrylamide gel phantoms for experimental tests. Albumin denatures and becomes opaque at temperatures above 58°C, enabling visual detection of lesions formed from denatured albumin. PSNE were vaporized using a 30-cycle, 3.2-MHz, at an acoustic power of 6.4 W (free-field intensity of 4,586 W/cm2) pulse from a single-element, focused high-power transducer. The vaporization pulse was immediately followed by a 15-s continuous wave, 3.2-MHz signal to induce ultrasound-mediated heating. Control experiments were conducted using an identical procedure without the vaporization pulse. Lesion formation was detected by acquiring video frames during sonication and post-processing the images for analysis. Broadband emissions from inertial cavitation (IC) were passively detected with a focused, 2-MHz transducer. Temperature measurements were acquired using a needle thermocouple.
Bubbles formed at the HIFU focus via PSNE vaporization enhanced HIFU-mediated heating. Broadband emissions detected during HIFU exposure coincided in time with measured accelerated heating, which suggested that IC played an important role in bubble-enhanced heating. In the presence of bubbles, the acoustic power required for the formation of a 9-mm3 lesion was reduced by 72% and the exposure time required for the onset of albumin denaturation was significantly reduced (by 4 s), provided that the PSNE volume fraction in the polyacrylamide gel was at least 0.008%.
The time or acoustic power required for lesion formation in gel phantoms was dramatically reduced by vaporizing PSNE into bubbles. These results suggest that PSNE may improve the efficiency of HIFU-mediated thermal ablation of solid tumors; thus, further investigation is warranted to determine whether bubble-enhanced HIFU may potentially become a viable option for cancer therapy.
High-intensity focused ultrasound (HIFU) is a medical procedure for the treatment of solid tumors [1–6]. In this procedure, ultrasound is focused into diseased tissue, and a fraction of the acoustic energy is converted into heat, primarily due to viscous absorption. Thermal ablation is possible by heating the tissue beyond the threshold temperature for protein denaturation. Using a focused transducer, the maximum point of heat deposition can be localized with millimeter precision. Thus, HIFU can be used to ablate solid tumors with minimal thermal damage to the surrounding and intervening tissue.
The focused ultrasound beam is normally generated with a single-element spherically focused transducer or by a transducer array. Because ultrasound can propagate through the tissue, the HIFU treatment does not require the insertion of probes and thus is noninvasive. In the past decade, HIFU has been used widely and has proven clinically to be successful in the treatment of a variety of cancers [7–9]. However, HIFU treatment of cancers that grow in organs behind by the rib cage or the skull (i.e., liver and brain cancers) is difficult because high attenuation of ultrasound in the bone increases the risk of thermal damage to the bone and the adjacent tissue. Additionally, acoustic intensity at the focus is reduced significantly, increasing the insonation time required for lesion formation. Therefore, a method that can reduce the acoustic power required for ablation and can increase the accuracy of the treatment while maintaining the therapeutic benefit will improve the clinical utility of HIFU for cancer therapy.
It has been well documented that the presence of microbubbles during sonication can increase the absorption of acoustic energy and can accelerate heating, which potentially could be used for increasing the efficiency of HIFU ablation [10–17]. While the results from documented studies are encouraging, microbubbles are not readily available in the tissue and thus must be created or introduced. Focused ultrasound can nucleate microbubbles in the tissue; however, it has been predicted that the applied rarefactional pressure must exceed 10 MPa in the absence of a nuclei [18, 19]. At such a high pressure, shock waves can form in the tissue, and the absorption of shock waves may heat the tissue beyond the boiling point in milliseconds . Because boiling bubbles can distort the lesion geometry significantly, the avoidance of boiling tissue is preferred during HIFU-mediated ablation [21, 22]. The introduction of exogenous agents can serve as the nuclei for cavitation in vivo, thus reducing the pressure threshold. Studies have demonstrated that systemically administered ultrasound contrast agents (UCAs) can nucleate cavitation for bubble-enhanced heating [14, 23–25]. However, the short lifespan of UCAs in circulation (<10 min) limits their use for bubble-enhanced thermal ablation [26, 27]. Furthermore, UCAs located in the blood vessels along the acoustic propagation path may attenuate HIFU significantly, resulting in unwanted heating and thermal damage in the healthy intervening tissue. The feasibility of using laser-illuminated gold nanoparticles has also been demonstrated for nucleating cavitation . However, this method was limited to superficial cancers due to lack of penetration of the laser in the tissue. Therefore, a reliable and consistent approach to nucleation locally within tumors is needed in order to take advantage of bubble-enhanced tumor ablation clinically.
In a previous study, we investigated the potential of using phase-shift nanoemulsions (PSNE) for nucleating microbubbles and reducing the pressure threshold for inertial cavitation (IC) . PSNE consist of nanodroplets composed of liquid perfluorocarbon, such as dodecafluoropentane (DDFP), and coated with phospholipids or albumin. The boiling point of DDFP in bulk is 29°C at standard atmospheric pressure, which is lower than physiological temperature (37°C). However, when liquefied DDFP is dispensed in the form of a nanoemulsion, DDFP is stable in the liquid phase at temperatures approaching 70°C. This is most likely due to the surface tension, which creates a pressure difference across the droplet interface that is inversely proportional to the droplet radius (i.e., Laplace pressure). It has been predicted that as the internal pressure is increased, the boiling point is elevated [30, 31]. In our previous study, we produced nanoemulsions with a mean diameter of 260 nm. Using the surface tension reported for naked perfluoropentane (PFP) droplets of 56 ± 1 mN/m , the Laplace pressure for our nanoemulsions was approximately 860 kPa. Nanoemulsions must be coated with surface-active molecules (i.e., surfactant) in order to inhibit fusion, and Rapoport et al. estimated that the coating may drop the surface tension for PFP to 30 mN/m . Using the Antoine equation log10P = A − B/(T + C)  and the Antoine constants A = 6.87362, B = 1,075.780, and C = 233.205 which were previously reported for n-pentane , we calculated a boiling point of 90.8°C for our nanoemulsions.
While stable at physiological temperature, the PSNE can be vaporized with short (<1 ms), high-amplitude (>5 MPa) acoustic pulses, a process known as acoustic droplet vaporization (ADV) . In a previous study, we used HIFU to vaporize PSNE dispersed throughout a polyacrylamide gel in a localized manner. The PSNE were vaporized only at the transducer focus when the rarefactional pressure exceeded a well-defined threshold (approximately 5 MPa); thus, HIFU provided a means to vaporize PSNE with exceptional spatial specificity and precision (i.e., on the order of millimeters). The PSNE reduced the pressure required for bubble formation in the gel phantoms and reduced the pressure for the onset of IC. Unlike UCA, scatter and attenuation from PSNE in liquid form are negligible; thus, unvaporized PSNE along the propagation path do not shield PSNE at the focus from the transmitted pulses. In addition, PSNE have no known toxicity, and DDFP has previously been tested clinically . Thus, PSNE is a good candidate for localized bubble nucleation in the tissue.
The main purpose of this study was to examine the feasibility of using vaporized PSNE to accelerate HIFU thermal lesion formation. The study was performed using PSNE dispersed throughout optically transparent albumin-containing gel phantoms, where lesion formation could be visualized and analyzed with video techniques. Furthermore, an ultrasound protocol was designed specifically for vaporizing the PSNE and driving bubble-enhanced heating in a localized manner. For the ultrasound exposure, a short (<1 ms), high-amplitude acoustic pulse was transmitted first to trigger acoustic droplet vaporization followed by continuous wave (CW) sonication. The acoustic intensity of the continuous wave exposure was below the ADV threshold; thus, we anticipate that the vaporization of additional PSNE will be avoided, limiting the impact of bubbles on HIFU ablation to the focal volume. In addition, in this study, DDFP droplets were coated with a phospholipid shell instead of albumin, in order to achieve narrower size distributions through extrusion. A narrower size distribution enables more reliable control over PSNE vaporization, and the reduced size would potentially increase the amount of PSNE accumulation in tumors in vivo through enhanced permeability and retention effect . We hypothesized that vaporized PSNE would significantly reduce the exposure time or acoustic power needed for lesion formation. Additionally, we explored the effect of inertial cavitation activity from vaporized PSNE on HIFU-mediated heating. Finally, the effect of PSNE concentration and acoustic intensity on final lesion shape and location was investigated.
Preparation of lipid-based phase-shift nanoemulsion
The phase-shift nanoemulsions consisted of DDFP (C5F12, CAS 678-26-2, Synquestlabs, Alachua, FL, USA) droplets, which were dispersed in saline and stabilized with a phospholipid monolayer shell. The shell components included 1,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC, Avanti Polar Lipids, Alabaster, AL, USA) and the lipopolymer 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethylene glycol)-2000] (DSPE-PEG2000, Avanti Polar Lipid) in the molar ratio of 25:1. DPPC worked as an emulsifier to stabilize the emulsion from coalescence, and DSPE-PEG2000 served as a polymer brush, limiting the interaction between droplets that could lead to fusion .
The nanoemulsions were prepared by combining ultrasound emulsification and pressure extrusion methods in order to get the desired mean diameter and size distribution. In the first step, 5.0 mg DPPC and 0.8 mg DSPE-PEG2000 were mixed with chloroform in a round-bottom flask. After mixing, the chloroform was removed by evaporation under vacuum, leaving a dry thin lipid film. The film was re-hydrated with 9.95 ml of saline to form a lipid solution. In the second step, 0.05 ml DDFP was added to the 9.95 ml of phospholipid saline solution and then emulsified with an ultrasonic liquid processor (Model VC505, Sonic & Materials, Newtown, CT, USA) for 60 s. The solution was kept in an ice water bath during sonication to avoid DDFP evaporation. The sonication step produced perfluorocarbon droplets with a lipid shell which were stable at physiological temperature. In the last step, the emulsion was forced 16 times at 20°C through a polycarbonate membrane with 200-nm pores (Whatman, Kent, ME, USA) using an extruder (10-ml LIPEX extruder, Northern Lipids, BC, Canada), yielding a 10-ml narrowly distributed suspension of DDFP nanodroplets. The nanoemulsion was stored in a sealed vial and refrigerated until further use. The size distribution of the nanoemulsion was determined at 37°C with a particle size analyzer (Model 90 Plus, Brookhaven Instruments, Holtsville, NY, USA).
Fabrication of tissue-mimicking phantom
All the tests of bubble-enhanced lesion formation were conducted in vitro using the albumin-containing acrylamide gel phantom originally developed by Lafon et al. . Slight modifications were made in order to uniformly distribute PSNE into the phantom and get better lesion visualization of lesion formation via video recording. The volume of each phantom was 2.65 × 2.65 × 1.71 cm3. This size was sufficient for this study since the largest lesion that was produced was 1 cm in length. The phantom was prepared by first mixing 2.1 ml of acrylamide (A9926, 40% 19:1 acrylamide/bis-acrylamide solution, Sigma-Aldrich Corporation, St. Louis, MO, USA), 1.2 ml of 1 M Tris buffer pH 8 (trizma hydrochloride and trizma base, Sigma-Aldrich Corporation), 0.1 ml of 10% (w/v) ammonium persulfate solution (APS, Sigma-Aldrich Corporation), and 1.08 g of bovine serum albumin (BSA, A3059, Sigma-Aldrich Corporation) in water. The entire solution was degassed for 1 h at 40°C, and then, PSNE was added. The mixture was stirred gently to get a uniform distribution of nanodroplets, and TEMED (87689, Sigma-Aldrich Corporation) was added last to initiate polymerization (1 μl TEMED/ml phantom solution). The phantom was submerged in a 12°C water bath during polymerization to remove the heat generated by the exothermic reaction. The volume fraction of DDFP in phantoms was used to describe the concentration of PSNE added, assuming that no DDFP was lost during emulsification and polymerization of the gels. Six different PSNE volume fractions were used for these experiments: 0.000%, 0.004%, 0.008%, 0.012%, 0.016%, and 0.020% (v/v), where 0.000% was used as the control case.
APS and TEMED in the recipe served as the cross linker and free radical generator, respectively. BSA served as an indicator of HIFU thermal ablation, as it denatures and turns white with sufficient heating [23, 39]. Since polyacrylamide gels are optically transparent, the denaturation of BSA was recorded real time using a hard disk drive camcorder with a 30-Hz frame rate (Everio, JVC, Yokohama, Japan). BSA also increases the attenuation coefficient of the gel. The speed of sound, density and attenuation of this type of phantom, as measured by Lafon et al.  at room temperature (22°C), were 1,044 ± 15 kg/m3, 1,544 ± 11 m/s, and 0.068 Np/cm at 3.2 MHz, respectively. All phantoms in this study were used on the same day of polymerization.
Power transducer calibration
The power transducer was a single-element spherically focused transducer with a 64-mm aperture and a 63-mm radius of curvature (Model H-102, Sonic Concepts, Woodinville, WA, USA). The power transducer was driven at its third harmonic (3.2 MHz), and the focal width and depth (pressure full width at half maximum (FWHM)) were 0.42 and 4.5 mm, respectively, at a temperature of 37°C. The excitation signal was provided by two function generators (33250A, Agilent, Santa Clara, CA, USA) in series with a 150-W RF amplifier (ENI A150, Rochester, NY, USA). As the desired waveforms in the tests were a short, high-amplitude pulse followed by CW exposure, two function generators were used. The first function generator delivered the ADV pulse and was triggered with the computer, while the second function generator was triggered by the first after the high-amplitude pulse was sent. A TTL delay circuit was used between the function generators to avoid overlap between the ADV pulse and CW exposure. The amplifier output impendence was matched to the transducer impedance via a matching network provided by the manufacturer. The acoustic power output of the transducer was calibrated with radiation force balance method as a function of the electrical input power , and the electrical power was measured with a power meter (E4419B, Agilent). The uncertainty of the measurements was 7%.
Passive cavitation detection
Ultrasound parameters: free-field acoustic powers and intensities to explore effect of PSNE vaporization on lesion formation
Acoustic power (corresponding intensity)
Initial pulse (30 cycle)
Continuous signal (15 s)
6.4 W (I = 4586 W/cm2)
0.8 W (I = 550 W/cm2)
0.8 W (I = 550 W/cm2)
0.8 W (I = 550 W/cm2)
2.7 W (I = 1957 W/cm2)
2.7 W (I = 1957 W/cm2)
Temperature measurement with thermocouple
A needle thermocouple (0.2 mm, Model HYP-0, Omega Engineering, Stamford, CT, USA) was used in some experiments to measure temperature elevations during ultrasound exposures with and without PSNE vaporization. The thermocouple was inserted into the phantom at the HIFU focal plane, parallel to the HIFU axis and 0.63 mm off axis laterally, where the acoustic pressure was reduced by 67% compared to the pressure at the focus. The thermocouple was placed out of the axis plane of the two transducers to avoid interference with PCD measurements. The thermocouple signal was amplified (Model SCXI-1112, National Instrument, Austin, TX, USA), digitized (Model PCI-6035E, National Instrument), and stored in the computer. The system was calibrated as described previously , and the accuracy was determined to be ±0.3°C. The alignment of the thermocouple to the HIFU beam was conducted in two steps. First, the needle thermocouple was inserted into the polyacrylamide gel under the guidance of a B-mode ultrasound (Terason 2000, Terason Ultrasound, Burlington, MA, USA). Second, the power transducer provided a pulsed sinusoid signal with 50% duty cycle and 1 Hz pulse repetition frequency. The phantom was moved until a maximum rate in the temperature rise during the HIFU on time was measured (dT < 5°C) [16, 42]. In the experiments, temperature elevation was measured using sonication parameters NVP1 and VP with a thermocouple inside a phantom mixed with 0.012% PSNE, and the corresponding PCD data were also recorded for comparison.
Monitoring lesion formation
All the tests were conducted with polyacrylamide gels separated into two groups. The first group was designed to test the feasibility of using vaporized PSNE to reduce the time or acoustic intensity required for lesion formation. The PSNE volume fraction was either 0.000% (sham), 0.008%, or 0.020%, and five tests were made for each volume fraction. In addition, four different acoustic intensities between 156 and 2,397 W/cm2 (with 7% uncertainty) were tested, with five tests at each intensity. All gel phantoms were sonicated with parameter VP (acoustic power of the tone burst exceeded vaporization threshold), NVP1 (acoustic power of the tone burst was below vaporization threshold), or NVP2 (acoustic power that is sufficient to cause lesion formation in the phantoms that did not contain PSNE). The second group was designed to investigate the effects of PSNE concentration in the lesion formation. Six different PSNE volume fractions were chosen (0.000%, 0.004%, 0.008%, 0.012%, 0.016%, 0.020%), and for each volume fraction, four or five sonications were made with parameters VP and NVP1. For all tests, the PCD and video data were recorded, stored, and processed later.
Image processing and statistics of lesion geometry
where N is the total number of pixels along the axis, r is the distance in pixels from the lesion border of that slice to its central axis, and C p is the length of each pixel in millimeters. It has been reported that the presence of cavitation or boiling will cause lesion distortion along the HIFU axis [21, 22]. By assuming that the lesion was divided into two parts with its middle point along the axis, the volumes of the proximal and distal parts of the lesion relative to the transducer were calculated as Vpre and Vpost, respectively. A distortion coefficient was then defined as Vpre/Vpost to represent the degree of distortion. If no distortion occurred, the lesion had a cigar shape and the coefficient was approximately equal to one. If distortion did occur, the lesion had a teardrop shape, resulting in a distortion coefficient greater than one.
Size distribution of nanoemulsions
Temperature and passive cavitation detection
Effect of vaporized PSNE on HIFU-mediated lesion formation
Effect of PSNE concentration on lesion formation
Effect of acoustic intensity on lesion formation with vaporized PSNE
While this study is a continuation of our research on nucleating bubbles with PSNE for the enhancement of ultrasound-mediated thermal ablation, we have made significant modifications to the protocol for making the nanoemulsions in order to produce them at a more optimal size for future in vivo applications. The addition of extrusion to the protocol narrowed the size distribution and reduced the mean diameter of the nanoemulsions (Figure 4). For this study, nanoemulsions were produced with a mean diameter below 200 nm. This is advantageous for in vivo applications since sizes below 200 nm are optimal for passive accumulation in tumors through the enhanced permeability and retention effect . Additionally, we coated the nanoemulsions with a mixture of phospholipids instead of albumin, which was used as the emulsifier in our previous study. Phospholipids can be conjugated with poly(ethylene glycol) (PEG), a polymer which has been shown to limit liposome aggregation and maintain a well-defined size distribution [46, 47]. More importantly, it has been shown that PEG increases the circulation time of systemically administered liposomes in vivo[48–50]. Thus, we speculate that coating our nanoemulsions with PEGylated lipids will increase the circulation time in vivo, and this is the subject of an ongoing study. PSNE have no known toxicity, and the components of PSNE (DDFP and PEGylated lipids) have previously been tested clinically [36, 51].
The primary objective of this study was to investigate the effect of vaporized PSNE on the time or acoustic power required for lesion formation. Studies were conducted with albumin-containing polyacrylamide gel phantoms because it allowed for control of the concentration of PSNE added as well as real-time visual observation of lesion formation. Variations in the lesion dimensions between experiments were observed with phantoms containing PSNE. Although care was taken during phantom preparation to ensure that the PSNE were well mixed within the acrylamide solutions, it was challenging to produce a perfectly uniform distribution of PSNE within the gel phantoms. Even tiny differences in the PSNE distribution within the phantoms can have a significant effect on the cavitation activity and thus heating rates, which can cause variation in lesion formation. For this reason, five experiments were performed for each condition tested. First, we confirmed that vaporized PSNE could nucleate inertial cavitation and accelerate HIFU-mediated heating within the phantom above the albumin denaturation temperature threshold. When PSNE were vaporized before CW exposure, the peak temperature measured outside the focal volume exceeded 70°C within 5 s (Figure 5). Provided the PSNE volume fraction was at least 0.008%, albumin denaturation was observed within 5 s in hydrogels treated after PSNE vaporization (Figure 9). More notably, it was possible to form a lesion of measurable volume (9 mm3) after PSNE vaporization (0.008% volume fraction) using 72% less power than the minimum required to denature albumin without vaporized PSNE (550 vs. 1,957 W/cm2, respectively). Furthermore, we found that the minimum acoustic intensity to denature albumin within 15 s is 89% less after PSNE vaporization (157 vs. 1,479 W/cm2). The reduction in acoustic power for lesion formation may have a significant impact on clinical applications of bubble-enhanced HIFU for cancer therapy, in particular for the treatment of brain and liver tumors. This reduction in power percentage (72%) to form measurable lesions exceeds the reduction in power percentage (30%) reported from a study of the effect of UCA on ultrasound-mediated thermal ablation in polyacrylamide gels . The difference in the reduction in power percentage is most likely due to attenuation of the transmitted acoustic waves in the gel by UCA prefocally. In our study, bubbles are not present along the beam path and the attenuation of PSNE is negligible compared to the polyacrylamide gel. Therefore, it is possible to localize the effect of bubbles on ultrasound-mediated heating and thermal ablation by vaporizing PSNE only at the transducer focus.
Lesions formed after PSNE vaporization had a predictable symmetric cigar shape at acoustic intensities between 157 and 550 W/cm2. However, at intensities greater than 830 W/cm2, the lesions formed a teardrop shape similar to other studies of lesions formed in gels and tissue in the presence of bubbles [12, 21, 52]. Based upon these observations, there may potentially be an optimal range of acoustic intensities that allow for taking advantage of bubble-enhanced heating while avoiding distortion in lesion shape. The symmetry and geometry of lesions formed at acoustic intensities of 157 W/cm2 (Figure 11) and 550 W/cm2 (Figure 7) were comparable to the lesions formed without PSNE vaporization. The symmetric cigar shape is advantageous for planning bubble-enhanced HIFU tumor ablation as it makes it possible to predict the lesion shape. However, it is important to note that lesions formed in the presence of vaporized PSNE still migrated towards the transducer, which must be accounted for in treatment planning. In addition to maintaining symmetry in lesion shape at low acoustic intensities (<550 W/cm2), the lesion volume was comparable in gels containing 0.008% and 0.020% PSNE. This was unexpected as several studies have reported an increase in the volume ablated by ultrasound in the presence of bubbles [24, 25, 53]. For example, Kaneko et al. reported that the volume of lesions formed in tumors after systemic administration of Levovist was 371 ± 104 mm3 compared with 166 ± 71 mm3 for saline . As an alternative to ultrasound contrast agents, Sokka et al. used a 0.5-s, 300-W tone burst to nucleate bubbles at the focus in a rabbit thigh . In our study, the lesion volume was primarily determined by the volume in which PSNE were vaporized. Thus, increasing the PSNE concentration above 0.008% had no detectable effect on lesion volume created at the aforementioned acoustic intensities.
Although the vaporized PSNE were localized to the HIFU focal plane, the lesions formed due to bubble-enhanced heating did tend to grow towards the transducer. Documented studies show that boiling may alter the lesion location due to the backscatter of incident waves by millimeter-sized bubbles [22, 55, 56]. Nonlinear wave propagation may also move the peak HIFU intensity towards the transducer, resulting in growth of the lesion towards the transducer . However, we did not observe any center shift for lesions generated without PSNE (Figure 7), which suggests that a shift in the lesion center most likely was not due solely to nonlinear wave propagation. While a prefocal shift in the location of maximum acoustic intensity due to nonlinear wave propagation may not be responsible for the migration of the lesion; it may shift the location of PSNE vaporization. Consequently, the impact of microbubbles on HIFU-mediated heating will be shifted towards the transducer, leading to the formation of lesions in the prefocal region. Another factor is that heating can induce changes in the acoustic attenuation of the gel phantom which could shift the focal region. Further studies on this topic are warranted as a sound fundamental understanding of the impact of cavitating bubbles on lesion location, size, and shape which is essential to the clinical translation of the technique for cancer therapy.
In conclusion, the feasibility of using PSNE to accelerate HIFU thermal lesion formation in albumin-containing gel phantoms was demonstrated. Lipid-coated phase-shift nanoemulsions were produced with a narrow size distribution (between 100 and 300 nm), which is important as the pressure threshold for vaporizing PSNE depends upon the droplet size. When driven to cavitate inertially, the bubbles formed by vaporizing PSNE reduced the acoustic intensity required for lesion formation in gel phantoms by as much as 89%. In addition, at an acoustic intensity of 550 W/cm2, the onset of lesion formation was reduced from 5 to 1 s of insonation. Furthermore, symmetrical lesions can be formed in the presence of bubbles provided that the acoustic intensity is kept low (<550 W/cm2). These results suggest that PSNE could eventually improve the efficiency of HIFU-mediated thermal ablation of solid tumors, thus potentially making bubble-enhanced HIFU a viable option for cancer therapy.
This work was supported financially by a BU/CIMIT Applied Healthcare Engineering Predoctoral Fellowship, a National Science Foundation Broadening Participation Research Initiation Grant in Engineering (BRIGE), and the National Institutes of Health (R21EB0094930).
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