Ultrasound-Responsive Cavitation Nuclei for Therapy and Drug Delivery

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.

The biological phenomena of sonoporation (i.e., membrane pore formation), stimulated endocytosis and opening of cell–cell contacts and the bio-effects of intracellular calcium transients, ROS generation, cell membrane potential change and cytoskeleton changes have been observed for several years (; ; ; ). However, other bio-effects induced by ultrasound-activated cavitation nuclei have recently been discovered. These include membrane blebbing as a recovery mechanism for reversible sonoporation (both for ultrasound-activated microbubbles [] and upon ADV []), extracellular vesicle formation (), suppression of efflux transporter P-glycoprotein (; ) and BBB (blood–brain barrier) transporter genes (). At the same time, more insight has been gained into the origin of the bio-effects, largely through the use of live cell microscopy. For sonoporation, real-time membrane pore opening and closure dynamics were revealed with pores <30 µm² closing within 1 min, while pores >100 µm² did not reseal () as well as immediate rupture of filamentary actin at the pore location () and correlation of intracellular ROS levels with the degree of sonoporation (). Real-time sonoporation and opening of cell–cell contacts in the same endothelial cells have been reported as well for a single example (). The applied acoustic pressure was found to determine uptake of model drugs via sonoporation or endocytosis in another study (). Electron microscopy revealed formation of transient membrane disruptions and permanent membrane structures, that is, caveolar endocytic vesicles, upon ultrasound and microbubble treatment (). A study by revealed that enhanced clathrin-mediated endocytosis and fluid-phase endocytosis occur through distinct signaling mechanisms upon ultrasound and microbubble treatment. The majority of these bio-effects have been observed in in vitro models using largely non-endothelial cells and may therefore not be directly relevant to in vivo tissue, where intravascular micron-sized cavitation nuclei will only have contact with endothelial cells and circulating blood cells. On the other hand, the mechanistic studies by , ) and do reveal translation from in vitro to in vivo. In these studies, ultrasound-activated microbubbles were found to induce a shear-dependent increase in intravascular adenosine triphosphate (ATP) from both endothelial cells and erythrocytes, an increase in intramuscular nitric oxide and downstream signaling through both nitric oxide and prostaglandins, which resulted in augmentation of muscle blood flow. Ultrasound settings were similar, namely, 1.3 MHz, mechanical index (MI) 1.3 for , ) and 1 MHz, MI 1.5 for , with MI defined as MI = P_–/ $\sqrt{f}$, where P_ is the derated peak negative pressure of the ultrasound wave (in MPa) and f the center frequency of the ultrasound wave (in MHz).

Whether or not there is a direct relationship between the type of microbubble oscillation and specific bio-effects remains to be elucidated, although more insight has been gained through ultrahigh-speed imaging of the microbubble behavior in conjunction with live cell microscopy. For example, there seems to be a microbubble excursion threshold above which sonoporation occurs (). further found that displacement of targeted microbubbles enhanced reversible sonoporation and preserved cell viability, whilst microbubbles that did not displace were identified as the main contributors to cell death.

All of the aforementioned biological observations, mechanisms and effects relate to eukaryotic cells. Study of the biological effects of cavitation on, for example, bacteria is in its infancy, but studies suggest that sonoporation can be achieved in Gram-negative bacteria, with dextran uptake and gene transfection being reported in Fusobacterium nucleatum (). More recent studies have investigated the effect of microbubbles and ultrasound on gene expression (; ; ). The findings are conflicting because although they all reveal a reduction in expression of genes involved in biofilm formation and resistance to antibiotics, an increase in expression of genes involved with dispersion and detachment of biofilms was also found (). This cavitation-mediated bio-effect needs further investigation.

Whilst there have been significant efforts to model the dynamics of ultrasound-driven microbubbles (; ), less attention has been paid to the interactions between microbubbles and cells or their impact upon drug transport. Currently there are no models that describe the interactions between microbubbles, cells and drug molecules. Several models have been proposed for the microbubble–cell interaction in sonoporation focusing on different aspects: cell expansion and microbubble jet velocity (), the shear stress exerted on the cell membrane (; ; ; ; ), microstreaming (), the shear stress exerted on the cell membrane in combination with microstreaming () or other flow phenomena (; ) generated by an oscillating microbubble. In contrast to the other models, propose that the microbubble-generated shear stress does not induce pore formation, but is instead due to microbubble fusion with the membrane and subsequent “pull out” of cell membrane lipid molecules by the oscillating microbubble. Models for pore formation (e.g., ) and resealing () in cell membranes have also been developed, but these models neglect the mechanism by which the pore is created. There is just one sonoporation dynamics model, developed by , that relates the uptake of the model drug propidium iodide (PI) to the size of the created membrane pore and the pore resealing time for a single cell in an in vitro setting. The model describes the intracellular fluorescence intensity of PI as a function of time, F(t), by

where α is the coefficient that relates the amount of PI molecules to the fluorescence intensity of PI-DNA and PI-RNA, D is the diffusion coefficient of PI, C₀ is the extracellular PI concentration, r₀ is the initial radius of the pore, β is the pore re-sealing coefficient and t is time. The coefficient α is determined by the sensitivity of the fluorescence imaging system, and if unknown, the equation can still be used because it is the pore size coefficient, α·πDC₀·r₀, that determines the initial slope of the PI uptake pattern and is the scaling factor for the exponential increase. A cell with a large pore will have a steep initial slope of PI uptake, and the maximum PI intensity quickly reaches the plateau value. A limitation of this model is that eqn (3) is based on 2-D free diffusion models, which holds for PI-RNA but not for PI-DNA because the latter is confined to the nucleus. The model is independent of cell type, as Fan et al. have reported agreement with experimental results in both kidney () and endothelial cells (). Other researchers have also used this model for endothelial cell studies and also classified the distribution of both the pore size and pore resealing coefficients using principal component analysis (PCA) to determine whether cells were reversibly or irreversibly sonoporated. In the context of BBB opening, have modeled the microbubble-generated shear and circumferential wall stress for 5-µm microvessels upon microbubble oscillation at a fixed MI of 0.134 for a range of frequencies (0.5, 1 and 1.5 MHz). The wall stresses were dependent upon microbubble size (range investigated: 2–18 µm in diameter) and ultrasound frequency. have also modelled the wall shear stress generated by microbubble (2 µm in diameter) destruction at 3 MHz for larger microvessels (200 µm in diameter). The presence of red blood cells was included in the model and was found to cause confinement of pressure and shear gradients to the vicinity of the microbubble. Advances in methods for imaging microbubble–cell interactions will facilitate the development of more sophisticated mechanistic models.