An international research team led by Dr. Tali Ilovitsh of the Biomedical Engineering Department at Tel Aviv University developed a noninvasive technology platform for gene delivery into breast cancer cells.
The technique combines ultrasound with tumor-targeted microbubbles. Once the ultrasound is activated, the microbubbles explode like smart and targeted warheads, creating holes in cancer cells’ membranes, enabling gene delivery.
Conducted over two years, the research was published on June 9 in the journal Proceedings of the National Academy of Sciences (PNAS).
Dr. Ilovitsh developed this breakthrough technology during her post-doctorate research at the lab of Prof. Katherine Ferrara at Stanford University.
The technique utilizes low frequency ultrasound (250 kHz) to detonate microscopic tumor-targeted bubbles.
In vivo, cell destruction reached 80% of tumor cells.
“Microbubbles are microscopic bubbles filled with gas, with a diameter as small as one tenth of a blood vessel,” Dr. Ilovitsh explains. “At certain frequencies and pressures, sound waves cause the microbubbles to act like balloons: they expand and contract periodically.
This process increases the transfer of substances from the blood vessels into the surrounding tissue. We discovered that using lower frequencies than those applied previously, microbubbles can significantly expand, until they explode violently.
We realized that this discovery could be used as a platform for cancer treatment and started to inject microbubbles into tumors directly.”
Dr. Ilovitsh and the rest of the team used tumor-targeted microbubbles that were attached to tumor cells’ membranes at the moment of the explosion, and injected them directly into tumors in a mouse model.
“About 80% of tumor cells were destroyed in the explosion, which was positive on its own,” says Dr. Ilovitsh. “The targeted treatment, which is safe and cost-effective, was able to destroy most of the tumor.
However, it is not enough. In order to prevent the remaining cancer cells to spread, we needed to destroy all of the tumor cells. That is why we injected an immunotherapy gene alongside the microbubbles, which acts as a Trojan horse, and signaled the immune system to attack the cancer cell.”
On its own, the gene cannot enter into the cancer cells. However, this gene aimed to enhance the immune system was co-injected together with the microbubbles.
Membrane pores were formed in the remaining 20% of the cancer cells that survived the initial explosion, allowing the entry of the gene into the cells. This triggered an immune response that destroyed the cancer cell.
“The majority of cancer cells were destroyed by the explosion, and the remaining cells consumed the immunotherapy gene through the holes that were created in their membranes,” Dr. Ilovitsh explains.
“The gene caused the cells to produce a substance that triggered the immune system to attack the cancer cell. In fact, our mice had tumors on both sides of their bodies. Despite the fact that we conducted the treatment only on one side, the immune system attacked the distant side as well.”
Dr. Ilovitsh says that in the future she intends to attempt using this technology as a noninvasive treatment for brain-related diseases such as brain tumors and other neurodegenerative conditions such as Alzheimer’s and Parkinson’s diseases.
“The blood-brain barrier does not allow for medications to penetrate through, but microbubbles can temporary open the barrier, enabling the arrival of the treatment to the target area without the need for an invasive surgical intervention.”
Cavitation nuclei for therapy
The most widely used cavitation nuclei are phospholipid-coated microbubbles with a gas core. For the 128 pre-clinical studies included in the treatment sections of this review, the commercially available and clinically approved Definity (Luminity in Europe; octafluoropropane gas core, phospholipid coating) (Definity 2011; Nolsøe and Lorentzen 2016) microbubbles were the most frequently used (in 22 studies).
Definity was used for studies on all applications discussed here, mostly for opening the BBB (12 studies). SonoVue (Lumason in the United States) is commercially available and clinically approved as well (sulfur hexafluoride gas core, phospholipid coating) (Lumason 2016; Nolsøe and Lorentzen 2016) and was used in a total of 14 studies for treatment of non-brain tumors (e.g., Xing et al. 2016), BBB opening (e.g., Goutal et al. 2018) and sonobactericide (e.g., Hu et al. 2018).
Other commercially available microbubbles were used that are not clinically approved, such as BR38 (Schneider et al. 2011) in the study by Wang et al. (2015d) and MicroMarker (VisualSonics) in the study by Theek et al. (2016).
Custom-made microbubbles are as diverse as their applications, with special characteristics tailored to enhance different therapeutic strategies. Different types of gasses were used as the core such as air (e.g., Eggen et al. 2014), nitrogen (e.g., Dixon et al. 2019), oxygen (e.g., Fix et al. 2018), octafluoropropane (e.g., Pandit et al. 2019), perfluorobutane (e.g., Dewitte et al. 2015), sulfur hexafluoride (Bae et al. 2016; Horsley et al. 2019) or a mixture of gases such as nitric oxide and octafluoropropane (Sutton et al. 2014) or sulfur hexafluoride and oxygen (McEwan et al. 2015).
While fluorinated gases improve the stability of phospholipid-coated microbubbles (Rossi et al. 2011), other gases can be loaded for therapeutic applications, such as oxygen for treatment of tumors (McEwan et al. 2015; Fix et al. 2018; Nesbitt et al. 2018) and nitric oxide (Kim et al. 2014; Sutton et al. 2014) and hydrogen gas (He et al. 2017) for treatment of cardiovascular disease.
The main phospholipid component of custom-made microbubbles is usually a phosphatidylcholine such as 1,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC), used in 13 studies (e.g., Dewitte et al. 2015; Bae et al. 2016; Chen et al. 2016; Fu et al. 2019), or 1,2-distearoyl-sn-glycero-3-phosphocholine (DSPC), used in 18 studies (e.g., Kilroy et al. 2014; Bioley et al. 2015; Dong et al. 2017; Goyal et al. 2017; Pandit et al. 2019). These phospholipids are popular because they are also the main components in Definity (Definity 2011) and SonoVue/Lumason (Lumason 2016), respectively.
Another key component of the microbubble coating is a polyethylene glycol (PEG)ylated emulsifier such as polyoxyethylene (40) stearate (PEG40-stearate; e.g., Kilroy et al. 2014) or the most frequently used 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-carboxy(polyethylene glycol) (DSPE–PEG2000; e.g., Belcik et al. 2017), which is added to inhibit coalescence and to increase the in vivo half-life (Ferrara et al. 2009).
In general, two methods are used to produce custom-made microbubbles: mechanical agitation (e.g., Ho et al. 2018) and probe sonication (e.g., Belcik et al. 2015). Both methods produce a population of microbubbles that is polydisperse in size.
Monodispersed microbubbles produced by microfluidics have recently been developed, and are starting to gain attention for pre-clinical therapeutic studies. Dixon et al. (2019) used monodisperse microbubbles to treat ischemic stroke.
Various therapeutic applications have inspired the development of novel cavitation nuclei, which is discussed in depth in the companion review by Stride et al. (2020). To improve drug delivery, therapeutics can be either co-administered with or loaded onto the microbubbles.
One strategy for loading is to create microbubbles stabilized by drug-containing polymeric nanoparticles around a gas core (Snipstad et al. 2017). Another strategy is to attach therapeutic molecules or liposomes to the outside of microbubbles, for example, by biotin–avidin coupling (Dewitte et al. 2015; McEwan et al. 2016; Nesbitt et al. 2018).
Echogenic liposomes can be loaded with different therapeutics or gases and have been studied for vascular drug delivery (Sutton et al. 2014), treatment of tumors (Choi et al. 2014) and sonothrombolysis (Shekhar et al. 2017). Acoustic Cluster Therapy (ACT) combines Sonazoid microbubbles with droplets that can be loaded with therapeutics for treatment of tumors (Kotopoulis et al. 2017).
The cationic microbubbles utilized in the treatment sections of this review were used mostly for vascular drug delivery, with genetic material loaded on the microbubble surface by charge coupling (e.g., Cao et al. 2015).
Besides phospholipids and nanoparticles, microbubbles can also be coated with denatured proteins such as albumin. Optison (Optison 2012) is a commercially available and clinically approved ultrasound contrast agent that is coated with human albumin and used in studies on treatment of non-brain tumors (Xiao et al. 2019), BBB opening (Kovacs et al. 2017b; Payne et al. 2017) and immunotherapy (Sta Maria et al. 2015).
Nano-sized particles cited in this review have been used as cavitation nuclei for treatment of tumors, such as nanodroplets (e.g., Cao et al. 2018) and nanocups (Myers et al. 2016); for BBB opening (nanodroplets; Wu et al. 2018); and for sonobactericide (nanodroplets; Guo et al. 2017a).
The physics of the interaction between bubbles or droplets and cells are described as these are the main cavitation nuclei used for drug delivery and therapy.
Physics of microbubble–cell interaction
Being filled with gas and/or vapor makes bubbles highly responsive to changes in pressure, and hence, exposure to ultrasound can cause rapid and dramatic changes in their volume.
These volume changes in turn give rise to an array of mechanical, thermal and chemical phenomena that can significantly influence the bubbles’ immediate environment and mediate therapeutic effects.
For the sake of simplicity, these phenomena are discussed in the context of a single bubble. It is important to note, however, that biological effects are typically produced by a population of bubbles and the influence of inter-bubble interactions should not be neglected.
A bubble in a liquid is subject to multiple competing influences: the driving pressure of the imposed ultrasound field; the hydrostatic pressure imposed by the surrounding liquid; the pressure of the gas and/or vapor inside the bubble; surface tension and the influence of any coating material; the inertia of the surrounding fluid; and damping caused by the viscosity of the surrounding fluid and/or coating, thermal conduction and/or acoustic radiation.
The motion of the bubble is determined primarily by the competition between the liquid inertia and the internal gas pressure. This competition can be characterized by using the Rayleigh–Plesset equation for bubble dynamics to compare the relative contributions of the terms describing inertia and pressure to the acceleration of the bubble wall (Flynn 1975a):
where R is the time-dependent bubble radius with initial value Ro, pG is the pressure of the gas inside the bubble, p∞ is the combined hydrostatic and time-varying pressure in the liquid, σ is the surface tension at the gas–liquid interface, ρL is the liquid density, IF is inertia factor and PF the pressure factor.
Flynn (1975a, 1975b) identified two scenarios: If the PF is dominant when the bubble approaches its minimum size, then the bubble will undergo sustained volume oscillations. If the inertia term is dominant (IF), then the bubble will undergo inertial collapse, similar to an empty cavity, after which it may rebound or it may disintegrate. Which of these scenarios occurs is dependent upon the bubble expansion ratio Rmax/Ro and, hence, the bubble size and the amplitude and frequency of the applied ultrasound field.
Both inertial and non-inertial bubble oscillations can give rise to multiple phenomena that affect the bubble’s immediate environment and hence are important for therapy. These include:
1. Direct impingement: Even at moderate amplitudes of oscillation, the acceleration of the bubble wall may be sufficient to impose significant forces on nearby surfaces, easily deforming fragile structures such as biological cell membranes (van Wamel et al. 2006; Kudo 2017) and blood vessel walls (Chen et al. 2011).
2. Ballistic motion: In addition to oscillating, the bubble may undergo translation as a result of the pressure gradient in the fluid generated by a propagating ultrasound wave (primary radiation force). Because of their high compressibility, bubbles may travel at significant velocities, sufficient to push them toward targets for improved local deposition of a drug (Dayton et al. 1999) or to penetrate biological tissue (Caskey et al. 2009; Bader et al. 2015; Acconcia et al. 2016).
3. Microstreaming: When a structure oscillates in a viscous fluid there will be a transfer of momentum as a result of interfacial friction. Any asymmetry in the oscillation will result in a net motion of that fluid in the immediate vicinity of the structure known as microstreaming (Kolb and Nyborg 1956). This motion will in turn impose shear stresses upon any nearby surfaces, as well as increase convection within the fluid. Because of the inherently non-linear nature of bubble oscillations (eqn ), both non-inertial and inertial cavitation can produce significant microstreaming, resulting in fluid velocities on the order of 1 mm/s (Pereno and Stride 2018). If the bubble is close to a surface then it will also exhibit non-spherical oscillations, which increases the asymmetry and hence the microstreaming even further (Nyborg 1958; Marmottant and Hilgenfeldt 2003).
4. Microjetting: Another phenomenon associated with non-spherical bubble oscillations near a surface is the generation of a liquid jet during bubble collapse. If there is sufficient asymmetry in the acceleration of the fluid on either side of the collapsing bubble, then the more rapidly moving fluid may deform the bubble into a toroidal shape, causing a high-velocity jet to be emitted on the opposite side. Microjetting has been reported to be capable of producing pitting even in highly resilient materials such as steel (Naudé and Ellis 1961; Benjamin and Ellis 1966). However, as both the direction and velocity of the jet are determined by the elastic properties of the nearby surface, its effects in biological tissue are more difficult to predict (Kudo and Kinoshita 2014). Nevertheless, as reported by Chen et al. (2011), in many cases a bubble will be sufficiently confined that microjetting will have an impact on surrounding structures regardless of jet direction.
5.Shock waves: An inertially collapsing cavity that results in supersonic bubble wall velocities creates a significant discontinuity in the pressure in the surrounding liquid leading to the emission of a shock wave, which may impose significant stresses on nearby structures.
6.Secondary radiation force: At smaller amplitudes of oscillation, a bubble will also generate a pressure wave in the surrounding fluid. If the bubble is adjacent to a surface, interaction between this wave and its reflection from the surface leads to a pressure gradient in the liquid and a secondary radiation force on the bubble. As with microjetting, the elastic properties of the boundary will determine the phase difference between the radiated and reflected waves and, hence, whether the bubbles move toward or away from the surface. Motion toward the surface may amplify the effects of phenomena 1, 3 and 6.
As described above, an oscillating microbubble will re-radiate energy from the incident ultrasound field in the form of a spherical pressure wave. In addition, the non-linear character of the microbubble oscillations will lead to the re-radiation of energy over a range of frequencies.
At moderate driving pressures, the bubble spectrum will contain integer multiples (harmonics) of the driving frequency; and at higher pressures, also fractional components (sub- and ultraharmonics).
In biological tissue, absorption of ultrasound increases with frequency and this non-linear behavior thus also increases the rate of heating (Hilgenfeldt et al. 2000; Holt and Roy 2001).
Bubbles will also dissipate energy as a result of viscous friction in the liquid and thermal conduction from the gas core, the temperature of which increases during compression. Which mechanism is dominant depends on the size of the bubble, the driving conditions and the viscosity of the medium.
Thermal damping is, however, typically negligible in biomedical applications of ultrasound as the time constant associated with heat transfer is much longer than the period of the microbubble oscillations (Prosperetti 1977).
The temperature rise produced in the surrounding tissue will be negligible compared with that occurring inside the bubble, especially during inertial collapse when it may reach several thousand Kelvin (Flint and Suslick 1991).
The gas pressure similarly increases significantly. Although only sustained for a very brief period, these extreme conditions can produce highly reactive chemical species, in particular reactive oxygen species (ROS), as well as the emission of electromagnetic radiation (sonoluminescence). ROS have been reported to play a significant role in multiple biological processes (Winterbourn 2008), and both ROS and sonoluminescence may affect drug activity (Rosenthal et al. 2004; Trachootham et al. 2009; Beguin et al. 2019).
Physics of droplet–cell interaction
Droplets consist of an encapsulated quantity of a volatile liquid, such as perfluorobutane (boiling point: –1.7°C) or perfluoropentane (boiling point: 29°C), which is in a superheated state at body temperature.
Superheated state means that although the volatile liquids have a boiling point below 37°C, these droplets remain in the liquid phase and do not exhibit spontaneous vaporization after injection. Vaporization can be achieved instead by exposure to ultrasound of significant amplitude via a process known as acoustic droplet vaporization (ADV) (Kripfgans et al. 2000).
Before vaporization, the droplets are typically one order of magnitude smaller than the emerging bubbles, and the perfluorocarbon is inert and biocompatible (Biro and Blais 1987). These properties enable a range of therapeutic possibilities (Sheeran and Dayton 2012; Lea-Banks et al. 2019).
For example, unlike microbubbles, small droplets may extravasate from the leaky vessels into tumor tissue because of the enhanced permeability and retention (EPR) effect (Long et al. 1978; Lammers et al. 2012; Maeda 2012), and then be turned into bubbles by ADV (Rapoport et al. 2009; Kopechek et al. 2013).
Loading the droplets with a drug enables local delivery (Rapoport et al. 2009) by way of ADV. The mechanism behind this is that the emerging bubbles give rise to similar radiation forces and microstreaming as described earlier in the Physics of the Microbubble–Cell Interaction. It should be noted that oxygen is taken up during bubble growth (Radhakrishnan et al. 2016), which could lead to hypoxia.
The physics of the droplet–cell interaction is largely governed by the ADV. In general, it has been observed that ADV is promoted by the following factors: large peak negative pressures (Kripfgans et al. 2000), usually obtained by strong focusing of the generated beam, high frequency of the emitted wave and a relatively long distance between the transducer and the droplet.
Another observation that has been made with micrometer-sized droplets is that vaporization often starts at a well-defined nucleation spot near the side of the droplet where the acoustic wave impinges (Shpak et al. 2014).
These facts can be explained by considering the two mechanisms that play a role in achieving a large peak negative pressure inside the droplet: acoustic focusing and non-linear ultrasound propagation (Shpak et al. 2016). In the following, lengths and sizes are related to the wavelength, that is, the distance traveled by a wave in one oscillation (e.g., a 1-MHz ultrasound wave that is traveling in water with a wave speed, c, of 1500 m/s has a wavelength, w (m), of c/f = 1500/106 = 0.0015, that is, 1.5 mm.
Because the speed of sound in perfluorocarbon liquids is significantly lower than that in water or tissue, refraction of the incident wave will occur at the interface between these fluids, and the spherical shape of the droplet will give rise to focusing.
The assessment of this focusing effect is not straightforward because the traditional way of describing these phenomena with rays that propagate along straight lines (the ray approach) holds only for objects that are much larger than the applied wavelength. In the current case, the frequency of a typical ultrasound wave used for insonification is in the order of 1–5 MHz, yielding wavelengths in the order of 1500–300 µm, while a droplet will be smaller by two to four orders of magnitude.
In addition, using the ray approach, the lower speed of sound in perfluorocarbon would yield a focal spot near the backside of the droplet, which is in contradiction to observations. The correct way to treat the focusing effect is to solve the full diffraction problem by decomposing the incident wave, the wave reflected by the droplet and the wave transmitted into the droplet into a series of spherical waves.
For each spherical wave, the spherical reflection and transmission coefficients can be derived. Superposition of all the spherical waves yields the pressure inside the droplet. Nevertheless, when this approach is only applied to an incident wave with the frequency that is emitted by the transducer, this will lead neither to the right nucleation spot nor to sufficient negative pressure for vaporization. Nanoscale droplets may be too small to make effective use of the focusing mechanism, and ADV is therefore less dependent on the frequency.
Treatment of tumors (non-brain)
The structure of tumor tissue varies significantly from that of healthy tissue which has important implications for its treatment. To support the continuous expansion of neoplastic cells, the formation of new vessels (i.e., angiogenesis) is needed (Junttila and de Sauvage 2013).
As such, a rapidly developed, poorly organized vasculature with enlarged vascular openings arises. Between these vessels, large avascular regions exist, which are characterized by a dense extracellular matrix, high interstitial pressure, low pH and hypoxia.
Moreover, a local immunosuppressive environment is formed, preventing possible anti-tumor activity by the immune system.
Notwithstanding the growing knowledge of the pathophysiology of tumors, treatment remains challenging. Chemotherapeutic drugs are typically administered to abolish the rapidly dividing cancer cells.
Yet, their cytotoxic effects are not limited to cancer cells, causing dose-limiting off-target effects. To overcome this hurdle, chemotherapeutics are often encapsulated in nano-sized carriers, that is, nanoparticles, that are designed to specifically diffuse through the large openings of tumor vasculature, while being excluded from healthy tissue by normal blood vessels (Lammers et al. 2012; Maeda 2012).
Despite being highly promising in pre-clinical studies, drug-containing nanoparticles have exhibited limited clinical success because of the vast heterogeneity in tumor vasculature (Barenholz 2012; Lammers et al. 2012; Wang et al. 2015d).
In addition, drug penetration into the deeper layers of the tumor can be constrained by high interstitial pressure and a dense extracellular matrix in the tumor. Furthermore, acidic and hypoxic regions limit the efficacy of radiation- and chemotherapy-based treatments because of biochemical effects (Mehta et al. 2012; McEwan et al. 2015; Fix et al. 2018).
Ultrasound-triggered microbubbles are able to alter the tumor environment locally, thereby improving drug delivery to tumors. These alterations are schematically represented in Figure 2 and include improving vascular permeability, modifying the tumor perfusion, reducing local hypoxia and overcoming the high interstitial pressure.
Several studies have found that ultrasound-driven microbubbles improved delivery of chemotherapeutic agents in tumors, which resulted in increased anti-tumor effects (Wang et al. 2015d; Snipstad et al. 2017; Zhang et al. 2018).
Moreover, several gene products could be effectively delivered to tumor cells via ultrasound-driven microbubbles, resulting in a downregulation of tumor-specific pathways and an inhibition in tumor growth (Kopechek et al. 2015; Zhou et al. 2015).
Theek et al. (2016) furthermore confirmed that nanoparticle accumulation can be achieved in tumors with low EPR effect. Drug transport and distribution through the dense tumor matrix and into regions with elevated interstitial pressure are often the limiting factors in peripheral tumors.
As a result, several reports have indicated that drug penetration into the tumor remained limited after sonoporation, which may impede the eradication of the entire tumor tissue (Eggen et al. 2014; Wang et al. 2015d; Wei et al. 2019).
Alternatively, microbubble cavitation can affect tumor perfusion, as vasoconstriction and even temporary vascular shutdown have been reported ex vivo (Keravnou et al. 2016) and in vivo (Hu et al. 2012; Goertz 2015; Yemane et al. 2018).
These effects were seen at higher ultrasound intensities (>1.5 MPa) and are believed to result from inertial cavitation leading to violent microbubble collapses. As blood supply is needed to maintain tumor growth, vascular disruption might form a different approach to cease tumor development.
Microbubble-induced microvascular damage was able to complement the direct effects of chemotherapeutics and antivascular drugs by secondary ischemia-mediated cytotoxicity, which led to tumor growth inhibition (Wang et al. 2015a; Ho et al. 2018; Yang et al. 2019b).
In addition, a synergistic effect between radiation therapy and ultrasound-stimulated microbubble treatment was observed, as radiation therapy also induces secondary cell death by endothelial apoptosis and vascular damage (Lai et al. 2016; Daecher et al. 2017). Nevertheless, several adverse effects have been reported because of excessive vascular disruption, including hemorrhage, tissue necrosis and the formation of thrombi (Goertz 2015; Wang et al. 2015d; Snipstad et al. 2017).
Furthermore, oxygen-containing microbubbles can provide a local oxygen supply to hypoxic areas, rendering oxygen-dependent treatments more effective. This is of interest for sonodynamic therapy, which is based on the production of cytotoxic ROS by a sonosensitizing agent upon activation by ultrasound in the presence of oxygen (McEwan et al. 2015, 2016; Nesbitt et al. 2018).
As ultrasound can be used to stimulate the release of oxygen from oxygen-carrying microbubbles while simultaneously activating a sonosensitizer, this approach has been reported to be particularly useful for the treatment of hypoxic tumor types (McEwan et al. 2015; Nesbitt et al. 2018). Additionally, low oxygenation promotes resistance to radiotherapy, which can be circumvented by a momentary supply of oxygen. Based on this notion, oxygen-carrying microbubbles were used to improve the outcome of radiotherapy in a rat fibrosarcoma model (Fix et al. 2018).
Finally, ultrasound-activated microbubbles promote convection and induce acoustic radiation forces. As such, closer contact with the tumor endothelium and an extended contact time can be obtained (Kilroy et al. 2014). Furthermore, these forces may counteract the elevated interstitial pressure present in tumors (Eggen et al. 2014; Lea-Banks et al. 2016; Xiao et al. 2019).
Apart from their ability to improve tumor uptake, microbubbles can be used as ultrasound-responsive drug carriers to reduce the off-target effects of chemotherapeutics. By loading the drugs or drug-containing nanoparticles directly into or onto the microbubbles, a spatial and temporal control of drug release can be obtained, thereby reducing exposure to other parts of the body (Yan et al. 2013; Snipstad et al. 2017).
Moreover, several studies have reported improved anti-cancer effects from treatment with drug-coupled microbubbles, compared with a co-administration approach (Burke et al. 2014; Snipstad et al. 2017).
Additionally, tumor neovasculature expresses specific surface receptors that can be targeted by specific ligands. Adding such targeting moieties to the surface of (drug-loaded) microbubbles improves site-targeted delivery and has been found to potentiate this effect further (Bae et al. 2016; Xing et al. 2016; Luo et al. 2017).
Phase-shifting droplets and gas-stabilizing solid agents (e.g., nanocups) have the unique ability to benefit from both EPR-mediated accumulation in the “leaky” parts of the tumor vasculature because of their small sizes, as well as from ultrasound-induced permeabilization of the tissue structure (Zhou 2015; Myers et al. 2016; Liu et al. 2018b; Zhang et al. 2018).
Several research groups have reported tumor regression after treatment with acoustically active droplets (Gupta et al. 2015; van Wamel et al. 2016; Cao et al. 2018; Liu et al. 2018b) or gas-stabilizing solid particles (Min et al. 2016; Myers et al. 2016). A different approach to the use of droplets for tumor treatment is ACT, which is based on microbubble-droplet clusters that upon ultrasound exposure, undergo a phase shift to create large bubbles that can transiently block capillaries (Sontum et al. 2015).
Although the mechanism behind the technique is not yet fully understood, studies have reported improved delivery and efficacy of paclitaxel and Abraxane in xenograft prostate tumor models (van Wamel et al. 2016; Kotopoulis et al. 2017). Another use of droplets for tumor treatment is enhanced high-intensity focused ultrasound (HIFU)-mediated heating of tumors (Kopechek et al. 2014).
Although microbubble-based drug delivery to solid tumors shows great promise, it also faces important challenges. The ultrasound parameters used in in vivo studies highly vary between research groups, and no consensus was found on the oscillation regime that is believed to be responsible for the observed effects (Wang et al. 2015d; Snipstad et al. 2017).
Moreover, longer ultrasound pulses and increased exposure times are usually applied in comparison to in vitro reports (Roovers et al. 2019c). This could promote additional effects such as microbubble clustering and microbubble translation, which could cause local damage to the surrounding tissue as well (Roovers et al. 2019a).
To elucidate these effects further, fundamental in vitro research remains important. Therefore, novel in vitro models that more accurately mimic the complexity of the in vivo tumor environment are currently being explored. Park et al. (2016) engineered a perfusable vessel-on-a-chip system and reported successful doxorubicin delivery to the endothelial cells lining this microvascular network.
While such microfluidic chips could be extremely useful to study the interactions of microbubbles with the endothelial cell barrier, special care of the material of the chambers should be taken to avoid ultrasound reflections and standing waves (Beekers et al. 2018).
Alternatively, 3-D tumor spheroids have been used to study the effects of ultrasound and microbubble-assisted drug delivery on penetration and therapeutic effect in a multicellular tumor model (Roovers et al. 2019b).
Apart from expanding the knowledge on microbubble–tissue interactions in detailed parametric studies in vitro, it will be crucial to obtain improved control over the microbubble behavior in vivo, and link this to the therapeutic effects. T
o this end, passive cavitation detection to monitor microbubble cavitation behavior in real time is currently under development, and could provide better insights in the future (Choi et al. 2014; Graham et al. 2014; Haworth et al. 2017).
Efforts are being committed to construction of custom-built delivery systems, which can be equipped with multiple transducers allowing drug delivery guided by ultrasound imaging and/or passive cavitation detection (Escoffre et al. 2013; Choi et al. 2014; Wang et al. 2015c; Paris et al. 2018).
More information: Tali Ilovitsh et al, Low-frequency ultrasound-mediated cytokine transfection enhances T cell recruitment at local and distant tumor sites, Proceedings of the National Academy of Sciences (2020). DOI: 10.1073/pnas.1914906117