Application of acoustic droplet vaporization in ultrasound therapy
© Zhou. 2015
Received: 22 June 2015
Accepted: 2 November 2015
Published: 11 November 2015
Microbubbles have been used widely both in the ultrasonic diagnosis to enhance the contrast of vasculature and in ultrasound therapy to increase the bioeffects induced by bubble cavitation. However, due to their large size, the lifetime of microbubbles in the circulation system is on the order of minutes, and they cannot penetrate through the endothelial gap to enter the tumor. In an acoustic field, liquefied gas nanoparticles may be able to change the state and become the gas form in a few cycles of exposure without significant heating effects. Such a phenomenon is called as acoustic droplet vaporization (ADV). This review is intended to introduce the emerging application of ADV. The physics and the theoretical model behind it are introduced for further understanding of the mechanisms. Current manufacturing approaches are provided, and their differences are compared. Based on the characteristic of phase shift, a variety of therapeutic applications have been carried out both in vitro and in vivo. The latest progress and interesting results of vessel occlusion, thermal ablation using high-intensity focused ultrasound (HIFU), localized drug delivery to the tumor and cerebral tissue through the blood-brain barrier, localized tissue erosion by histotripsy are summarized. ADV may be able to overcome some limitations of microbubble-mediated ultrasound therapy and provide a novel drug and molecular targeting carrier. More investigation will help progress this technology forward for clinical translation.
Ultrasound has been used in clinics since 1940s, mostly for diagnostic and physiotherapy purposes at low intensity. With the advance of electrics, computer technology, and transducer design in the 1980s, moderate or high-intensity ultrasound was applied in medicine with promising results. Since then, therapeutic ultrasound is emerging as an interesting and important research topic not only for physicists and engineers but also for medical practitioners. Cavitation is one of the important mechanisms in ultrasound therapy [1, 2]. However, the distribution of bubble nuclei in the soft tissue and blood is quite sparse. In order to induce the desired bioeffect in the acoustic field using low energy in vivo, artificial bubble nuclei can be introduced for cavitation. Ultrasound contrast agents (UCAs), such as gaseous microbubbles, have been used in sonographic diagnosis in cardiology clinics as cavitation enhancers in some therapeutic applications, such as blood clot fragmentation, tissue erosion, drug delivery, gene transfection, and necrosis formation in the thermal ablation [3–7]. However, UCAs have relatively large size (1–10 μm) and short in vivo circulation lifetime (i.e. a few minutes). Thus, re-administration of UCAs requires repeated sonication. Furthermore, these microbubbles cannot extravasate into tumor tissue for efficient targeting of agents to deep within the tumor.
The phenomenon of acoustic droplet vaporization (ADV), in which the phase shift of the core from liquid to gas triggered by an acoustic wave, was been described in the 1990s and was used extensively in imaging for preclinical trials [8–10]. The physics are based on the vapor pressure of the liquid, which is a function of temperature, and not necessarily based upon the liquid chemistry. In theory, ADV could be employed with any liquid that has a normal boiling point near or below the body temperature. Fluorocarbons are good candidates, particularly the perfluorocarbons (PFCs) because they have low solubility in aqueous formulations and relatively low toxicity. Perfluoropentane (PFC5) is perhaps the most commonly used in ADV because of its favorable transition temperature (~29 °C), its good combination of high vapor pressure, low solubility in blood, price, and availability. Due to the low solubility and diffusivity of PFC gases in water, bubbles can remain stable in an aqueous solution much longer than air bubbles in the same size . Clearance of PFC nanoparticles occurs not through metabolism but slow dissolution into the surrounding medium. The resulting half-life in the tissue is much longer than that of microbubble, ranging from 4 days (i.e., perfluorooctylbromide) to 65 days (i.e., perfluorotripropylamine). Since these liquid droplets were much smaller than gas bubbles, they could traverse the lungs and provide contrast in the left heart by ADV better than the Food and Drug Administration (FDA)-approved contrast agent Albunex in the diagnosis of ventricular opacification [12, 13]. Phase-change colloids for ultrasonic imaging resulted in the advent of EchoGen™ in 1996. It shows that PFC nanodroplets (~200 nm in diameter) can be vaporized with short ultrasound pulses emitted from clinical diagnostic machines. However, the diagnostic use declined rapidly with the advent of other imaging agents. Some PFCs exhibit echogenicity due to a lower acoustic impedance than water, but not as significant as that of microbubbles. Sonography with the aid of UCAs has been proved to be a sensitive and inexpensive imaging technique in cardiovascular and oncological diagnosis. Currently, this technology has been approved by the FDA for the echocardiographic examination of wall motion abnormalities and ventricular contraction. So, the use of ADV for standard clinical imaging of cardiovascular organs and systems may not experience much resurgence in the future.
Capillaries in the diameter between 5 and 10 μm have only a single layer of endothelial cells. Although capillary vasculature in normal tissues has tight inter-endothelial junctions and therefore inhibits extravasation of nanoparticles, a variety of tumors have a porous and disorganized defective microvasculature with pore cutoff size ranging between 380 and 780 nm, which permits the so called enhanced permeability and retention (EPR) effect . This characteristic allows submicron vesicles to penetrate tumor’s capillaries in a passive way. In addition to the EPR, tumors also have poor lymphatic drainage, which further ensures prolonged circulation time of nanoparticles. To avoid both extravasation to normal tissues and recognition by cells of the reticulo-endothelial system (RES), the nanoparticles are usually coated with polyethylene glycol (PEG) chains to suppress blood protein adsorption.
Microdroplets encapsulated in albumin or lipid shells in ADV are used to generate embolization and show promise for spatially and temporally targeted and substantial tissue occlusion. Both intracardiac and intravenous injections repeatedly could produce ADV in chosen arteries (i.e., renal or segmental arteries) as seen by sonography. Localized cortex occlusion was achieved with a maximum regional flow reduction of >90 % and an average organ perfusion reduction of >70 % using intracardiac injections, and vaporization from intravenous injections resulted in a substantial echogenicity increase with an average half-life of 8 min per droplet dose, which could result in the onset of cell death and possible tumor treatment via ischemic necrosis [15, 16]. Ninety percent of the pefluoropentane (PFP) concentration in the blood can be eliminated from the body within 10 min after their administration, which diminishes the risk of toxicity .
For nanosized droplets, surface tension effects dominate the largest expansion. Therefore, a sharp decline in the expansion factor is found in the lower nanometer range of droplets. In order to utilize the EPR effect for enhanced intratumoral diffusion, droplets of 100–200 nm in size is preferred. However, these droplets have the expansion factor of 2.5–3.5, resulting in bubbles on the order of 250–700 nm. Although the vaporized bubbles still have cavitation effects for enhanced thermal therapy or drug delivery, they are much smaller than preferable for diagnostic imaging (on the order of 1–5 μm). Post-extravasation droplet/bubble coalescence may increase echogenicity, but no in vivo evidence has confirmed such hypothesis. Therefore, there exists a design trade-off in echogenicity and droplet diffusivity . In comparison, a transition occurs at the low micrometer range because of the dominant effect of the lowest ambient pressure, approximately 25 times the original diameter for frequencies between 1.5 and 8 MHz at the threshold from 4.5 to 0.75 MPa peak rarefactional pressure, respectively. This agent might be useful for tissue occlusion in cancer treatment, as well as for phase aberration corrections in acoustic imaging . But, its penetration to interstitial space is limited.
Acoustic activation with significant temporal and spatial specificity has led to a diverse set of in vitro and in vivo applications, not only strictly for intravascular administration because liquid PFC emulsions have a broad range of sizes. In the last decade, ADV was proposed and explored for embolic occlusion therapy, molecular imaging, drug delivery, aberration correction, high-intensity focused ultrasound (HIFU) sensitization, and tissue erosion by histotripsy [8, 16, 19, 20]. In this review paper, the physics of acoustic droplet vaporization is first introduced to understand this phenomenon. The manufacture processes are summarized briefly. Then, several medical applications are listed and followed by the discussion of the future development.
Acoustic cavitation is one possible mechanism of vaporization. However, the ADV threshold is found less than the inertial cavitation (IC) threshold. Thus, an IC event is not necessary to cause vaporization, though IC nuclei, such as ultrasound contrast agent and polystyrene microspheres, external to the droplet can lower the ADV threshold. ADV occurs first and may provide a bubble then going through IC. Because the IC threshold remained constant with different gas saturation of the bulk fluid and the droplet diameter, it is likely that the nucleus for IC is not external to the droplet and may be the ADV bubble itself. In addition, it is possible to achieve ADV with or without IC . Meanwhile, the intensity required to produce higher harmonics (the onset of stable cavitation) using the Fourier transform to the acoustic emission signals measured by passive cavitation detection (PCD) was about 10-fold greater than the occurrence of the fundamental frequency (phase shift). The acoustic intensity required for the baseline shift (the onset of IC) was about sixfold higher than that required for the higher harmonic frequencies . Larger droplet requires less acoustic intensity to produce IC. PFC6 nanoemulsions at 476 kHz required higher intensity for phase shift compared to PFC5 while opposite conclusions were found at 20 kHz .
The behavior of vaporized submicron droplets immediately following vaporization is different to their micronsized counterparts. The efficiency of droplet vaporization is dependent on the initial droplet nucleation rate, and the stability of the newly created bubbles following vaporization is determined by recondensation into the liquid form due to an increase in Laplace pressure. There is a threshold bubble radius of approximately 800 nm (in concordance with the theoretical prediction using the Homogeneous Nucleation Theory) above which the bubbles are stable against recondensation and below which uncoated bubbles are prone to recondensation. Long interval time does not significantly affect the threshold pressure for recondensation. The efficiency of initial droplet nucleation is quite high (on the order of at least 10 %), as approximately ten droplets may be sufficient for the production of a detectable amount of bubbles with comparable echogenicity as the single microbubble in sonography. However, up to 90 % of the newly formed bubbles may recondense into the liquid, resulting in the efficiency of stable bubble generation dropping to even below 1 %. Following vaporization and bubble formation, the newly created bubbles retain their initial coating material and undergo rapid expansion (up to 10 μm in diameter) followed by stable oscillations for a number of cycles eventually coming to rest or rapid shrinkage or collapse to a radius of less than 1 μm. During the initial growth and collapse, the bubbles may undergo fragmentation and coalescence. There is no dependence of bubble coalescence rate on ultrasound excitation pressure. The probability of bubble survival is significantly higher for coalesced bubbles than the noncoalesced ones following vaporization because of increased total volume and size of the bubble as well as the reduced Laplace pressure, but decreased with the excitation acoustic pressure. Presence of a shell reduces the effective surface tension on the bubble and subsequently the Laplace pressure. Bubbles with higher concentration of surfactant coating are more likely to avoid condensation than those with less surfactant. Coalescence of bubbles may also increase the surfactant concentration on the final combined bubble. Uncoated PFP bubbles above the radius of approximately 1 μm do not recondense but dissolve on the order of milliseconds .
Bubble expansion may not follow the prediction of the ideal gas law because of the influx of dissolved gases from the surrounding medium. The influx of dissolved oxygen into a gas phase inside the emulsion occurs during the rarefactional phase of the acoustic wave, followed by the rectified diffusion of dissolved gases into the formed bubble. The solubility of gas increases with pressure according to Henry’s law. Although some gases may be released during the acoustic compression phase, its degree is less than the uptake due to higher gas solubility under pressure and a smaller surface area of the compressed bubbles. When the acoustic wave is terminated, equilibrium corresponding to the ambient pressure is restored; gases with super-equilibrium concentrations diffuse out, thus restoring nanodroplets. The noncondensable gas will not condense along with the PFC and may not completely dissolve into the surrounding liquid, leaving a very small bubble of noncondensable gas that easily nucleates in the next cycle of PFC boiling, leading to an even larger bubble, and subsequently more diffusion of noncondensable gas into the bubble. The vapor pressure of water is 3.17 kPa at 25 °C, an order lower than that of a small PFC (i.e., 29.1 kPa for PFC6). Thus, PFC bubbles will form before homogeneous nucleation of water vapor bubbles. However, in practice, a small portion of water evaporates into the first bubbles to form. The first perfluoropentane bubble to form contains 3.5 and 4.4 % water vapor in equilibrium at 25 and 37 °C, respectively. Perfluorohexane has a lower vapor pressure, so the corresponding values are slightly higher, 9.8 and 11.6 % at 25 and 37 °C, respectively . The complexity in the vapor components would lead to complicated bubble dynamics after vaporization.
The performance of droplets in the acoustic field is determined by many parameters, such as ultrasound pulses, droplet properties, and environmental variables. A viscous fluid, limited tissue space, and high elasticity of surrounding tissue may retard the initial expansion of the gas bubble generated by ADV and re-condense the gas nucleus at lower rarefactional pressures. Greater acoustic amplitude will start bubble growth sooner and have more total time for growth. Similarly, low-frequency ultrasound provides a longer time window for growth. For example, using a 2-cycle sinusoid pulse at the ultrasound frequency of 1 MHz, bubble formation was found at rarefactional pressure threshold of 1 MPa, while at 8 MHz the threshold reached approximately 3 MPa for decafluorobutane nanoemulsions with peak diameter of 200 nm . The ADV threshold was shown to be inversely proportional with degree of superheating and independent of pulse length at clinically relevant ultrasound frequencies, concentrations of droplet, and the volume fractions of the PFC, which may be useful for in vivo application because of limited information of droplets on site [25, 35, 36]. For longer pulse lengths (i.e., milliseconds), vaporization can be induced at a decreased ultrasound pressure. In comparison, for microsized droplets (90 % < 6 μm), the vaporization drops with the ultrasound frequency, from 4.5 to 0.75 MPa peak rarefactional pressure between 1.5 and 8 MHz . It takes a longer time for the microdroplets to completely vaporize than the submicron ones. However, a larger emulsion grows into a larger bubble, which is due to a longer time of subpressurization and a larger subpressurization driving force by the smaller Laplace pressure and the initiation of gas evolution earlier in the acoustic cycle.
Perfluorocarbon (PFC) including perfluorohexane (C6F14), dodecafluoropentane (DDFP, C5F12), and decafluorobutane (DFB, C4F10) is a common material for liquid emulsion formulation. The use of others, such as phosphatidylethanolamine (PE), soybean phosphatidylcholine (PC), pefluoropentane (PFP), perfluorohexane (PFH), perfluorobutane (PFB), perfluoro-15-crown-5-ether (PFCE), perfluorooctylbromide (PFOB, C8F17Br), and perfluorotripropylamine was also mentioned [21, 37–40]. Some manufacture approaches are listed below.
The first method is similar to that of microbubble manufacture. Liquefied PFC gas and degassed, deionized water were mixed and emulsified with an ultrasonic liquid processor for about 30 s. Sometimes, coarse emulsification by a vortex mixer or an amalgamator was performed before the use of sonicator. The resulting emulsion was poured slowly into albumin or lipid (i.e., phospholipids) or polymer solution or surfactant to coat the droplet, and the shell was used to stabilize the emulsion by lowering the surface tension as well as inhibiting coalescence. These uncoated emulsions are stable for 30–60 min and then start to coalesce after 2–3 h, which may be due to the nature of hydrocarbon. The surfactant (i.e., PF68) can interact with phospholipid shells and increase the surface area of each droplet . Subsequently, the size of the emulsions decreases with higher surfactant ratio. But, drug release from the inner core may be much slower than that from the interfacial shell . Double emulsion (i.e., oil-in-PFC-in-water) could also be manufactured using similar method . In addition, micelles and liposomes could also be produced depending on the lipid concentration. This method is simple in manufacture. In order to have uniform size distribution, the emulsion needs to pass through filter paper with certain pore size, which may reduce the manufacturing efficiency.
Droplet could be formulated via extrusion. PFC was first condensed in a secure container over dry ice, poured into a glass vial, crimped, and stored at −20 °C to preserve the liquid state. The lipid solution was cooled to approximately −2 to −5 °C in order to avoid freezing of the aqueous solution and then mixed with liquid PFC. The mixture was extruded by several passes through porous membrane filter. After that, the emulsion was stored in a crimped vial at 4 °C with room air in the headspace . Both the submicrometer and the micrometer size (as large as 12–15 μm) droplets were made via extrusion.
Microbubble condensation method induced by pressurization and low temperature allows simple production of high-yield nanoemulsions from volatile compounds . Polydisperse distribution of microbubbles was first produced using standard mechanical agitator, cooled to room temperature before being immersed in a CO2/isopropanol bath at −5 to −10 °C, and swirled gently for about 1 min. Headspace pressure in the vial was then increased by an adjustable air-pressure source to revert the emulsions to the liquid state. Pressurization and temperature condensation of microbubbles are effective and advantageous in producing ADV nanoemulsions in comparison to extrusion and emulsion-based methods. The condensed droplet size has a peak of 200–300 nm, which corresponding well with the prediction using the ideal gas law from the original bubbles in the 1–2 μm range. Variability in the droplet size (some samples having content below 100 nm while others having only content greater than 200 nm) may be due to both insistencies in applied pressure and temperature at the time of condensation and formation of micelles and liposomes as a function of lipid concentration. Some content as large as 2–4 μm was also found in the sample despite of low percentage (1.5–4 %), and the upper size limit seems to increase with lipid concentration. The samples with a high number of viable outputs could be obtained using a similar technique as traditional microbubble preparation .
A microfluidic channel with a flow-focusing structure can also be used for uniform nanoemulsion production by forcing a central stream of a dispersed phase and two side sheath flows of a continuous phase through a small orifice. The size of droplets including the shell thickness is well controlled and adjusted by the geometry of microfluidic channel and flow rate. The minimum size is the width of the orifice. A high capillary number will lead to droplet generation in the jetting mode, and breaking off the tip of the dispersed phase finger due to Rayleigh capillary instability will result in polydisperse production with a polydispersity index less than 5 % . In addition, over 2 weeks, the mean droplet diameter decreased less than 4 % from 4.5 ± 0.2 μm to 4.3 ± 0.3 μm. However, manufacture speed of this method is not very high. Multiplexing numerous flow-focusing circuits would scale the throughput. A 10× scale-up would significantly reduce the production time of 1 × 109 droplets on the order of minutes similar to the amount and time formed by mechanical agitation.
Vessel occlusion (embolotherapy)
It is suggested to “starve” cells to death by restricting their blood supply from the feeder vessel . One method of treating tumors or other malformations is to occlude the blood flow to them with gas bubbles, which can effectively shrink its size. By using focused ultrasound, vaporization can be induced intentionally and accurately in the feeder arteries of tumors, resulting in the occlusion with a high degree of spatial specificity [15, 16]. For frequencies between 1.5 and 8 MHz, the threshold of peak rarefactional pressure for vaporization decreases from 4.5 to 0.75 MPa for microdroplets (90 % < 6-μm diameter) . Large gas bubbles (>30 μm) could be formed temporarily. Vessel occlusion via ADV has been explored in rodents and dogs (76-μm mean length and 36-μm mean diameter in capillary and 25-μm mean length and 11-μm mean diameter in feeder vessel) and may be translated to clinical use soon . Image-based hyper-echogenicity from ADV of intra-articular (IA) and intravenous (IV) injections after sonic exposure (9.2-MPa peak negative pressure, 3.5-MHz frequency, 13 cycles, pulse repetition frequency of 1 kHz and ISPTA of 10 W/cm2) was monitored for approximately 90 min, and cortex perfusion was reduced by >60 % of its original value for more than 1 h, which could be long enough for the onset of cell death and possible tumor treatment via ischemic necrosis. However, in these studies, the control kidney on the contralateral side also showed 18 % of the decrease in regional blood flow relative to the preocclusion baseline, which may be due to the balance of the urinary output between the treated and untreated kidneys . IV administration results in a lower gas bubble yield, which may be due to the filtering in the lung, dilution in the blood volume, or other circulatory effects. So, large volume of the droplet is required for IV administration in comparison to IA injection (0.3 vs. 0.03 mL) . However, a raft of recent data shows that starving of a tumor can lead to increased epithelial-to-mesenchymal transition and metastatic escape leading to worse overall survival in patients [47, 48]. Thus, the future of this approach is unclear.
In comparison to the other vascular occlusion modalities, ADV could improve both diagnostic and therapeutic ultrasound fields. ADV-based angiography can provide sensitive feedback on the effect of ultrasonic therapy in models of pancreatic cancer, breast cancer, and kidney function. Despite clinical competition from other modes of vascular occlusion, ADV-induced embolotherapy has great potential in vital organ tissues (brain, liver, eye, etc.), in which revascularization after therapeutic healing is desired. Occlusion can also be accompanied by deposition of a chemotherapy drug to increase local cytotoxic effects and minimize systemic effects . For example, thrombin release from PFC emulsion could extend the duration of ADV-generated microbubble occlusions. Angiostatin and endostatin suppress the development of supply arteries for tumor growth .
ADV-generated occlusion also has the potential to instigate hemostasis for vascular damage or internal bleeding. Vaporizing in the capillaries could cause the rupture of these vessels and subsequent red blood cell extravasation. The encapsulated chemical embolic agent, such as thrombin, within a PFC emulsion could be released locally and noninvasively upon ADV with precision on the order of millimeters and with no need for ionizing fluoroscopy as in transcatheter embolization for sustained embolization. Furthermore, prolonged ischemia generated by vascular occlusion may activate water soluble, bioreductive prodrugs, such as NLCQ-1 , encapsulated within the emulsion .
Occlusion is also beneficial in the thermal ablation induced by radio frequency, microwave, or high-intensity focused ultrasound (HIFU). The blood flow in the tumor acts as a heat sink, dissipates the heat via vascular cooling, and subsequently reduces the efficacy of the treatment. Reduced blood flow by occlusion could also induce hypoxia in tumors, but such occlusion prevents drugs diffusion into the target.
Overall, embolotherapy must be carried out carefully because many arterial emboli could create infarcts in the heart or brain or travel to distant vascular bed where they could cause unwanted arterial occlusion, ischemia, and potentially infarction . The performance of ADV-based occlusion is determined by the dynamics of emulsion and bubbles in the transport (i.e., interaction with vessel bifurcations downstream) and lodging of the generated microbubbles in the microvasculature near the site of vaporization (i.e., sliding along the vascular space). If the droplets have stealth character, the generated gas bubbles will easily coalesce into sufficiently large one(s) to occlude arterioles and capillaries. However, vascular occlusion may not work well for superficial tumors. It was found that reduction in blood flow lowers the core temperature of superficial tumors, which in turn increases the survival of tumor cells by approximately two orders of magnitude [52, 53]. Moreover, significant growth delays were found to require occlusion times on the order of 4 h, which requires repetitive treatments of ADV.
In the acoustic field, bubble shielding effects are of importance, in which the resident bubbles will reflect or back scatter the energy of incoming acoustic pulses toward the source. As a result, it is possible to vaporize additional nanoemulsions easily in the prefocal region during HIFU tumor ablation, which could lead to unpredictable prefocal lesion formation . In HIFU ablation, PFCs with a higher boiling point, such as PFH, may be more appropriate. Inertial cavitation may significantly reduce the ADV threshold for HIFU exposures longer than a millisecond. Vaporizing albumin-coated DDFP microdroplets by heat alone required temperatures as much as 40 °C above the boiling point , and such discrepancy may increase as the droplet size decreases. In addition, vaporized microbubbles may be manipulated to enhance targeting through acoustic radiation force [54–56].
Targeted tumor chemotherapy is an active research area with great potential. The concept of a “magic bullet,” a drug carrier that responds to a certain stimulus, was first proposed by Ehrlich in the early twentieth century. However, most drug delivery techniques do not have the capability of temporally and spatially specific targeting as ultrasound. The control of localized delivery is especially important for drugs which possess narrow therapeutic windows and will allow the deleterious effects on healthy tissues to be minimized. Ultrasound also provides the ability of target diagnosis and focusing onto deeply located tissues. Drug-loaded microbubbles under ultrasound exposure can potentially target drugs to specific sites. Therapeutic agents are typically incorporated into the microbubbles by attachment to or insertion in the shell, complexation of secondary carriers to the microbubble shell, or incorporation within a fluid inside the shell. Chemotherapeutic drug (i.e., paclitaxel, PTX) was tightly retained by nanodroplets stabilized with poly(ethylene oxide)-co-polycaprolactone (PEG-PCL) block copolymer. But, it was effectively released into tumor after acoustic radiation, which resulted in effective tumor regression [38, 57]. The interaction of ultrasonic pulse with a payload containing microbubbles involves a number of mechanisms, such as acoustic cavitation, heating, radiation forces, and sonoporation . Stable bubble cavitation generates strong shear stress close to the bubble surface, sufficient to shear cell membranes. Inertial cavitation produces shock waves and high-speed microjets, which also disrupt cell membranes. The transient increase in cell membrane permeability allows the uptake of drugs, genes, and peptides from a variety of carriers (polymeric micelles, liposomes, and nanoemulsions).
In the past decades, extensive studies have been carried out to understand the mechanism of ADV, simulate the bubble dynamics in the acoustic field, and apply this phenomenon in medical investigation. Due to the complexity involved in this event, including nucleation, vaporization, oscillation, recondensation, and dissolution, and a large range of droplet size used in the study (both submicron and micron), more effort is required to fully understand the mechanisms. Some insights have already been obtained owing to the use of ultrafast photography with a large number of frames. Meanwhile, in order to enhance its clinical performance and safety for quick transition and wide acceptance of the applications, some aspects, but not limited, may have more interests for continuous investigation in the following days.
Inclusion of paramagnetic nanoparticles, fluorescent nanoparticles, and radioisotopes enables the detection of PFC emulsions in 1H MR T1-weighted imaging, SPECT-CT, and optical fluorescence imaging [70–73]. Fluorine NMR has a large chemical shift range (~300 ppm) allowing the examination of multiple agents simultaneously with minimal signal overlap. The integral of quantitative NMR signal, such as the fluorine atoms within the core of PFC nanodroplets, is proportional to the amount of particles being interrogated for MR molecular imaging. The 19F isotope of fluorine has a natural abundance of near 100 %, but virtually zero in the biological presence. Therefore, the chemical shift of the PFCs can be easily differentiated from the background. The nanoparticles for molecular imaging must have a long circulation time, highly sensitive and selective binding to the epitope of interest, prominent contrast-to-noise enhancement, acceptable toxicity, ease of clinical use, and compatibility with commercially available imaging systems. Multifunctional activity can be realized by incorporating more targeting ligands, imaging agents, and drugs into the formulation simultaneously. Materials can be covalently or noncovalently linked to the particle surface, dissolved in the coating, or carried in particle interiors for cellular deposition and activation. Multivalent interactions between high-avidity agents and cell surface may partially overcome the dissociation of the targeting ligand itself in the nanomolar range .
Most of the ADV techniques mentioned above are still in preclinical studies but have potential for clinical use in specialty applications. Overall, ADV has a bright future because the small size of nanodroplets greatly reduces the rate of clearance compared to larger ultrasound contrast agent bubbles and yet provides the advantages of ultrasonographic contrast, acoustic cavitation, and acceptable toxicity of conventional perfluorocarbon contrast agent bubbles. Injection of fluorocarbon emulsions induces short- and long-term effects that spontaneously resolve within 12–24 h. The adverse effect of “pulmonary hyperinflation” observed upon administration to humans is less likely than that for rabbits, pigs, or monkeys. In addition, toxicity of surfactants that stabilize the droplets must also be considered. Natural phospholipids, polysaccharides, and human proteins may be the best stabilizing agents to use.
The capability of simultaneous imaging and targeted drug delivery of nanoparticles shows great promise for individualizing therapeutics and could enable conclusive assurance that the drug is reaching the intended target in a much higher effective drug concentration. Imaging can also be used to monitor distribution of released drugs. For example, liposome colabeled with gadolinium (Gd) could allow MRI monitoring. Fluorescent markers on the drug can be detected using bioluminescent imaging system. Such information allows optimal timing of external stimulus application. Furthermore, approval by regulatory agencies may be very slow although the FDA has specific pathways for combined devices that go through two separate departments . This regulatory obstacle may temper the enthusiasm of pharmaceutical companies to pursue development, given the expense and risk of clinical trials . ADV is a complex process and not a fully understood phenomenon, involving many variables that are not always straightforward to regulate for safety and may be some variables not taken into consideration yet.
New treatment strategies will be developed, and operation parameters will be optimized in the following years to enhance the outcome. The internalization of targeted or nontargeted droplets into cells and subsequent vaporization could lead to new tissue-specific therapy. A co-injection of both nano- and microdroplets could simultaneously release drugs in the tumor interstitium by extravasated nanodroplets and occlude the vascular space by larger ones . If the lowest vaporization pressure is required, the PFC with the lowest boiling point allows for stable circulation at physiological temperatures will be used. The optimal particle size of the emulsion should result in a vascular persistence that is clinically efficacious while limiting deep-tissue retention. Droplets may remain stable until a desired “activation pulse” is delivered to create desired bioeffects. More measurements are required to establish the acoustic thresholds for gas expansion as a function of acoustic parameters (frequency, amplitude, pulse length, etc.) and the characteristics of the droplets (chemical composition, size, stabilizing surfactants, temperature, etc.) .
Acoustic droplet vaporization, a phenomenon of phase shift from liquid to gas in the acoustic field, is a new approach of producing bubbles to enhance the bioeffects of ultrasound therapy with high accuracy of spatial and temporal control. The burst produced from commercially available echosonographic systems is strong enough to induce ADV for microdroplets. Submicron size of droplets allows penetration into tumors via EPR effects and much prolonged half-life time in the circulating system, overcoming the limitation of the commonly used ultrasound contrast agents (i.e., microbubbles), but needs more acoustic pressure to vaporize. Its recent applications, such as vessel occlusion, thermal ablation, drug delivery, and histotripsy, have already shown very promising results. Although most in vivo studies are still in the preclinical stage, translation into clinics may occur in the following years. In comparison to bubble cavitation of microbubbles, that of vaporized droplets has more stable cavitation and less inertial cavitation, which makes ADV good candidate for localized drug delivery with the minimal damage to the surrounding tissue. In order to understand the mechanism more clearly, effort in the theoretical modeling and experimental observation is highly desired.
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