**1. Introduction**

212 Non-Viral Gene Therapy

Liu Y, Miyoshi H, Nakamura M. Encapsulated ultrasound microbubbles: therapeutic application in drug/gene delivery. J Control Release 2006; 114: 89–99. Litzinger DC, Brown JM, WalaI, Kaufman SA, et al. Fate of cationic liposomes and their

Marmottant P, Hilgenfeldt S. Controlled vesicle deformation and lysis by single oscillating

Marshall E. Gene Therapy Death Prompts Review of Adenovirus Vector.

Mehier-Humbert S, Bettinger T, Yan F, Guy RH. Ultrasound- mediated gene delivery: kinetics of plasmid internalization and gene expression. J Control Release 2005; 104: 203–211. Michel MS, Erben P, Trojan L, Schaaf A, Kiknavelidze K, Knoll T et al. Acoustic energy: a

Miller MW, Miller DL , Brayman AB. A review of in vitro bioeffects of inertial ultrasonic

Muruve DA, Barnes MJ, Stillman IE, Libermann TA. Adenoviral gene therapy leads to rapid

Newman CMH, Bettinger T. Gene therapy progress and prospects: Ultrasound for gene

Rahim AA, Taylor SL, Bush NL, TerHaar GR, Bamber JC, Porter CD. Spatial and acoustic

Shohet RV, Chen S, Zhou Y-T, Wang Z, Meidell RS, Unger RH, Grayburn PA.

TerHaar G. Therapeutic applications of ultrasound . Prog Biophys Mol Biol 2007; 93: 111–129. Thomas CE, EhrhardtA, Kay MA. progress and problems withthe use of viral vectors for gene therapy. 2003;4: www.nature.com/reviews/genetics p346-358. Unger, E C, Mccreery Tp, Sweitzer Rh. Ultrasound Enhances Gene Expression of Liposomal

Zhigang W, Zhiyu L, Haitao R, Hong R, Qunxia Z, Ailong H et al. Ultrasound-mediated

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benign kidney cells. Anticancer Res 2004; 24: 2303–2308.

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vivo. Hum Gene Ther.1999;(10): 965–976.

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complex with oligonucleotive in vivo. Biochimica et Biophysica Acta (BBA) -

bubbles. Nature 2003; 423: 153–156. 12 Sundaram J, Mellein BR, Mitragotri S. An experimental and theoretical analysis of ultrasound-induced permeabilization of cell membranes. Biophys J 2003; 84: 3087–3101. 13 Brujan EA. The role of cavitation microjets in the therapeutic applications of ultrasound. Ultrasound Med Biol 2004;

new transfection method for cancer of the prostate, cancer of the bladder and

cavitation from a mechanistic perspective. Ultrasound in Med. & Biol.. Vol. 22. No.

induction of multiple chemokines and acute neutrophil dependent hepatic injury in

pressure dependence of microbubble- mediated gene delivery targeted using

Echocardiographic Destruction of Albumin Microbubbles Directs Gene Delivery to

microbubble destruction enhances VEGF gene delivery to the infarcted

transfection activity of HPMA-stabilized DNA polyplexes with prolonged plasma

Human gene therapy holds great promise in treating not only hereditary genetic disorders, but also disease states such as cancer and viral infections, and contingencies such as stroke or myocardial infarctions. It can be achieved by delivery of a correct gene into target cells with genetic deficiency or mutations, or by transfer of a therapeutic agent such as agents targeting a cancer-causing oncogene, growth factor gene, antisense oligonucleotides (ODN), or small interfering RNA (siRNA) to correct the disease state using either viral or nonviral vectors. Viral gene therapy has succeeded in many animal disease models {Snyder 1999}, and has progressed to clinical trials {Hacein-Bey-Abina *et al.* 2002; Kay *et al.* 2000}. However, significant obstacles remain, including immune responses {Manno *et al.* 2006} or tumor genesis {Hacein-Bey-Abina *et al.* 2003}. A nonviral approach would provide a safer strategy. The potential for therapeutic ultrasound (US) to effect minimally invasive nonviral gene transfer has long been recognized, and a growing body of evidence indicates that significant enhancement of transgene expression can be achieved by using high frequency acoustic energy. In addition to its well-known role in providing inexpensive, real-time imaging capability, US has been used therapeutically for years {Herzog *et al.* 1999}. The most common therapeutic application involves low acoustic intensities and is intended to heat deep tissues; *e.g*., as used in sports medicine. At the other 'end' of the acoustic intensity spectrum is HIFU (high intensity focused ultrasound), which can be used to ablate {Fischer *et al.* 2010} or to liquefy tissues {Hall *et al.* 2009}. US of intermediate intensities has been applied to many systems, together with exogenous microbubbles [MBs], to use the acoustically-forced behavior of the MBs to generate desired bioeffects. The latter usually involves changing the permeability of endogenous barriers to otherwise impermeable materials (*e.g*., drugs or macromolecules). Many gene therapies have been attempted by direct intramuscular or intraparenchymal injection of gene vectors; these vectors gain immediate access to the interstitial space and must then traverse the plasma membrane of the targeted cells. US contrast agents are almost always administered intravascularly. When accompanied by a gene vector, the first barrier encountered is the vascular endothelium. The next are other vascular anatomical features (*e.g.*, the basement membrane, smooth muscle layer, *etc.*) and then the outer cell membrane of the cells one hopes to target. Finally, DNA needs to be transferred across the nuclear membrane to enter the nucleus for efficient gene expression.

This review will focus almost entirely on the use of ultrasound targeted microbubble destruction (UTMD) as a means by which to deliver foreign DNA (or drugs or photo

Ultrasound-Mediated Gene Delivery 215

contingency, although *in vitro* UTMD transfection techniques are being used for cell-based therapies {Otani *et al.* 2009}. Specific applications to various organs or for various clinical conditions will be discussed in **§5**. Here it is sufficient to note that in UTMD-based gene delivery studies, much of the work has focused on model or surrogate systems; *e.g*., the delivery of reporter genes rather than therapeutic ones. However, therapeutic gene transfer

Therapeutic US has the potential for enhancing minimally invasive gene therapies. For gene therapies involving naked DNA vectors in particular, UTMD techniques have many desired characteristics. These include (1) low toxicity of all components of the treatment, (2) low immunogenicity of the vectors, (3) low invasiveness (*e.g*., the vector and gas bodies can be administered intravascularly, and for sonographically-accessible organs, the therapeutic US can be applied through the skin), (4) there is good potential for repeated application, (5) organs can be targeted with high specificity, and (6) the technique has broad applicability (again related to sonographic accessibility). However, low efficiencies remain a problem.

Here we hope to provide a sense of the types of MBs often used in US-mediated gene therapies, the ways in which US can cause these MBs to be destroyed (either gradually or abruptly) and/or otherwise activated, and the mechanisms by which UTMD-induced microvascular damage, extravasation, and target cell uptake of gene vectors may occur. A

In broad terms, the targeting 'part' of UTMD therapies is based *principally* on the fact that MBs present in tissues respond dynamically only if the area is exposed to US. Ligands incorporated into the MB shell may enhance accumulation of the MBs and any vector load they carry in a region of interest, but it is the selective acoustic exposure which 'activates' the MBs. As mentioned, site-specific ligands can be added to MB shells to improve retention in regions of interest. This field has received considerable attention {Ferrara *et al.* 2009}. Some degree of targeting can also be achieved by acoustically 'pushing' bubbles to a region of interest. A propagating acoustic field exerts a radiation force on MBs {Sarvazyan *et al.* 2010}, and this force can propel them in the direction of wave propagation. If the MB is undergoing radial oscillations {Emmer *et al.* 2007}, the bubble lurches forward, slowing as it expands and accelerating as it collapses. Radiation force can be used to concentrate MBs

more comprehensive discussion of the physics can be found in {Wu & Nyborg 2008}.

along a specific wall of an intracavitary space {Horie *et al.* 2010}, or blood vessel.

interstitial space. Cellular permeabilization is also associated with MB activation.

**2.2 Ultrasound contrast agents and other stabilized gas bodies** 

The MB destruction 'part' of UTMD occurs in response to the acoustic exposure, and may result in release of lipid or aqueous-phase drugs {Smith *et al.* 2010} or produce small-scale damage to microvessels which allows normally impermeable materials (drugs, plasmids, even objects as large as cells or MB fragments) to escape the vascular lumen and enter the

Microbubble US contrast agent evolution from early, agitated saline or sugar solutions to protein-shelled agents containing air to lipid- or polymer-shelled agents using relatively insoluble gases was rapid. The evolution of these agents continues {Qin *et al.* 2009}, as do new applications for them {Cosgrove & Harvey 2009}. US contrast agents are typically micron-sized gas bodies which are stabilized against diffusion by a shell {Overvelde *et al.* 2010; Sarkar *et al.* 2009}; most contrast agent MBs have mean diameters of ~2 µm. New

effects such as tumor volume reduction have been reported in some model systems.

**2.1 What is ultrasound targeted microbubble destruction?** 

reactive nanoparticles, as other examples for which there are numerous publications) into targeted host tissues and cells. The literature on this topic is growing at an incredible rate, because almost immediately following the commercial availability of microbubble-based US contrast agents in the 1980s, it was recognized that acoustic cavitation could cause potentially undesirable bioeffects, desirable ones, or both.

Exploitation of acoustically-activated MBs for therapeutic effect remains an exciting topic. Indeed, harnessing the dynamical behavior of acoustic MBs is somewhat of the 'holy grail' of acoustically-targeted gene (or drug) delivery {Lindner 2009}. However, a healthy skepticism has been (see, *e.g*., {Villanueva 2009} and should continue to be, applied to claims of great successes achieved using UTMD, as replication studies are few, and contradictory findings not unusual. Skepticism should obtain especially when considering 'black box' US studies.

A reasonably coherent picture of US-mediated gene therapy is emerging. With UTMD techniques, the issue ultimately reduces to the fact that US can force dynamic behaviors of MBs. Endogenous MBs are absent in most tissues of the body {Carstensen *et al.* 2000; Gross *et al.* 1985}. Exogenously-administered MBs are generally inert without acoustic exposure, but can be made to pulsate gently, or to undergo violent but highly localized dynamic behavior when driven by acoustic fields. Targeting can thus be achieved by methods as simple as co-administering MBs and gene vectors and exposing the targeted tissue to US, which will thus be the only site where bubble activation occurs. As we shall see, bioeffects typically arise when MBs are driven at pressure amplitudes sufficient to produce nonlinear bubble oscillations.

Here we will discuss the issues broadly, in the hope that the reader will gain a general understanding of the techniques, applications, apparent mechanisms, and some insights into what has been achieved. We focus most of our discussion on *in vivo* studies, as US-assisted gene therapy *in vivo* continues to be a more challenging problem than US-enhanced cell permeabilization *in vitro* or even in intact *ex vivo* tissues (see; *e.g*. {Kodama *et al.* 2005}). We have striven to be as jargon-free as possible, and have neglected mathematical treatments of the topics discussed here, as these can be found elsewhere. A few simple **abbreviations** are used: **US** (ultrasound); **MB** (microbubble); **Pa** or **Pr** (acoustic pressure amplitude or rarefaction pressure, respectively); and **pDNA** (plasmid DNA). Some words on US exposure metrics are also necessary. In some applications of therapeutic US (*e.g*., tissue heating), the acoustic intensity in dimensions of W/cm2 is the parameter of interest. In contrast, the occurrence and character of acoustic cavitation *in vivo* is largely determined by the presence or absence of exogenous MBs and by *peak acoustic intensity, or more properly, by the peak acoustic pressures*. These are expressed in units of mega- (MPa) or kiloPascals (kPa), where 1 MPa = 10 atm. Acoustic intensity scales as the square of the pressure amplitude. The quality of acoustic reporting in the US-mediated gene therapy literature varies widely. We will describe most acoustic exposures in terms of the pressure amplitude (Pa) or the peak rarefactional acoustic pressure (Pr), sometimes inferred by us. In any case, it is not our purpose here to be rigorously quantitative.

#### **2. Ultrasound-targeted microbubble destruction: Physics & technology**

There are many potential applications for UTMD gene delivery; a large body of work has been conducted using both *in vitro* and *in vivo* model systems to understand if, and how, UTMD 'works' to produce therapeutic effects. Most studies have as their ultimate goal the application of the technology to effect minimally-invasive treatment of disease or

reactive nanoparticles, as other examples for which there are numerous publications) into targeted host tissues and cells. The literature on this topic is growing at an incredible rate, because almost immediately following the commercial availability of microbubble-based US contrast agents in the 1980s, it was recognized that acoustic cavitation could cause

Exploitation of acoustically-activated MBs for therapeutic effect remains an exciting topic. Indeed, harnessing the dynamical behavior of acoustic MBs is somewhat of the 'holy grail' of acoustically-targeted gene (or drug) delivery {Lindner 2009}. However, a healthy skepticism has been (see, *e.g*., {Villanueva 2009} and should continue to be, applied to claims of great successes achieved using UTMD, as replication studies are few, and contradictory findings not unusual. Skepticism should obtain especially when considering 'black box' US studies. A reasonably coherent picture of US-mediated gene therapy is emerging. With UTMD techniques, the issue ultimately reduces to the fact that US can force dynamic behaviors of MBs. Endogenous MBs are absent in most tissues of the body {Carstensen *et al.* 2000; Gross *et al.* 1985}. Exogenously-administered MBs are generally inert without acoustic exposure, but can be made to pulsate gently, or to undergo violent but highly localized dynamic behavior when driven by acoustic fields. Targeting can thus be achieved by methods as simple as co-administering MBs and gene vectors and exposing the targeted tissue to US, which will thus be the only site where bubble activation occurs. As we shall see, bioeffects typically arise when MBs are driven at pressure amplitudes sufficient to produce nonlinear

Here we will discuss the issues broadly, in the hope that the reader will gain a general understanding of the techniques, applications, apparent mechanisms, and some insights into what has been achieved. We focus most of our discussion on *in vivo* studies, as US-assisted gene therapy *in vivo* continues to be a more challenging problem than US-enhanced cell permeabilization *in vitro* or even in intact *ex vivo* tissues (see; *e.g*. {Kodama *et al.* 2005}). We have striven to be as jargon-free as possible, and have neglected mathematical treatments of the topics discussed here, as these can be found elsewhere. A few simple **abbreviations** are used: **US** (ultrasound); **MB** (microbubble); **Pa** or **Pr** (acoustic pressure amplitude or rarefaction pressure, respectively); and **pDNA** (plasmid DNA). Some words on US exposure metrics are also necessary. In some applications of therapeutic US (*e.g*., tissue heating), the acoustic intensity in dimensions of W/cm2 is the parameter of interest. In contrast, the occurrence and character of acoustic cavitation *in vivo* is largely determined by the presence or absence of exogenous MBs and by *peak acoustic intensity, or more properly, by the peak acoustic pressures*. These are expressed in units of mega- (MPa) or kiloPascals (kPa), where 1 MPa = 10 atm. Acoustic intensity scales as the square of the pressure amplitude. The quality of acoustic reporting in the US-mediated gene therapy literature varies widely. We will describe most acoustic exposures in terms of the pressure amplitude (Pa) or the peak rarefactional acoustic pressure (Pr), sometimes inferred by us. In any case, it is not our

**2. Ultrasound-targeted microbubble destruction: Physics & technology** 

There are many potential applications for UTMD gene delivery; a large body of work has been conducted using both *in vitro* and *in vivo* model systems to understand if, and how, UTMD 'works' to produce therapeutic effects. Most studies have as their ultimate goal the application of the technology to effect minimally-invasive treatment of disease or

potentially undesirable bioeffects, desirable ones, or both.

bubble oscillations.

purpose here to be rigorously quantitative.

contingency, although *in vitro* UTMD transfection techniques are being used for cell-based therapies {Otani *et al.* 2009}. Specific applications to various organs or for various clinical conditions will be discussed in **§5**. Here it is sufficient to note that in UTMD-based gene delivery studies, much of the work has focused on model or surrogate systems; *e.g*., the delivery of reporter genes rather than therapeutic ones. However, therapeutic gene transfer effects such as tumor volume reduction have been reported in some model systems.

Therapeutic US has the potential for enhancing minimally invasive gene therapies. For gene therapies involving naked DNA vectors in particular, UTMD techniques have many desired characteristics. These include (1) low toxicity of all components of the treatment, (2) low immunogenicity of the vectors, (3) low invasiveness (*e.g*., the vector and gas bodies can be administered intravascularly, and for sonographically-accessible organs, the therapeutic US can be applied through the skin), (4) there is good potential for repeated application, (5) organs can be targeted with high specificity, and (6) the technique has broad applicability (again related to sonographic accessibility). However, low efficiencies remain a problem.

#### **2.1 What is ultrasound targeted microbubble destruction?**

Here we hope to provide a sense of the types of MBs often used in US-mediated gene therapies, the ways in which US can cause these MBs to be destroyed (either gradually or abruptly) and/or otherwise activated, and the mechanisms by which UTMD-induced microvascular damage, extravasation, and target cell uptake of gene vectors may occur. A more comprehensive discussion of the physics can be found in {Wu & Nyborg 2008}.

In broad terms, the targeting 'part' of UTMD therapies is based *principally* on the fact that MBs present in tissues respond dynamically only if the area is exposed to US. Ligands incorporated into the MB shell may enhance accumulation of the MBs and any vector load they carry in a region of interest, but it is the selective acoustic exposure which 'activates' the MBs. As mentioned, site-specific ligands can be added to MB shells to improve retention in regions of interest. This field has received considerable attention {Ferrara *et al.* 2009}. Some degree of targeting can also be achieved by acoustically 'pushing' bubbles to a region of interest. A propagating acoustic field exerts a radiation force on MBs {Sarvazyan *et al.* 2010}, and this force can propel them in the direction of wave propagation. If the MB is undergoing radial oscillations {Emmer *et al.* 2007}, the bubble lurches forward, slowing as it expands and accelerating as it collapses. Radiation force can be used to concentrate MBs along a specific wall of an intracavitary space {Horie *et al.* 2010}, or blood vessel.

The MB destruction 'part' of UTMD occurs in response to the acoustic exposure, and may result in release of lipid or aqueous-phase drugs {Smith *et al.* 2010} or produce small-scale damage to microvessels which allows normally impermeable materials (drugs, plasmids, even objects as large as cells or MB fragments) to escape the vascular lumen and enter the interstitial space. Cellular permeabilization is also associated with MB activation.

#### **2.2 Ultrasound contrast agents and other stabilized gas bodies**

Microbubble US contrast agent evolution from early, agitated saline or sugar solutions to protein-shelled agents containing air to lipid- or polymer-shelled agents using relatively insoluble gases was rapid. The evolution of these agents continues {Qin *et al.* 2009}, as do new applications for them {Cosgrove & Harvey 2009}. US contrast agents are typically micron-sized gas bodies which are stabilized against diffusion by a shell {Overvelde *et al.* 2010; Sarkar *et al.* 2009}; most contrast agent MBs have mean diameters of ~2 µm. New

Ultrasound-Mediated Gene Delivery 217

**2.4 Microbubble dynamics: An overview of how gas bodies respond to ultrasound**  At very low Pa, MB volume oscillations are related linearly to Pa. At modest Pas of a few hundred kPa (or less), MB volume oscillations become non-linearly related to Pa, with bubble expansion being relatively slow and bubble collapse much faster, being governed by the inertia of the in-rushing surrounding fluid to a greater extent than by the compressive phase of the applied pressure field; hence the phenomenon of bubbles undergoing acoustically-driven expansions followed by rapid, inertially-dominated collapse is termed 'inertial cavitation'. This typically occurs when a MB has been driven by the rarefaction phase of the acoustic wave to a diameter roughly 2 – 3 times the initial diameter {Chomas *et al.* 2001}. At very modest Pas, inertial cavitation can be stable and sustained for an almost indefinite number of acoustic cycles. For example, {Church & Carstensen 2001} found that surfactant-coated Sonazoid MBs could undergo repetitive inertial collapses and rebounds when driven by 2.5 MHz US at acoustic pressures greater than about 0.3 – 0.4 MPa; for this agent, irreversible post-collapse fragmentation occurred at a Pa of ~ 1.5 MPa. Others have reported that a 3 µm bubble exposed to 2.25 MHz US can be expected to undergo expansion, inertial collapse, and fragmentation at a Pa just over 0.3 MPa {Chomas *et al.* 2001}. The albumin-shelled, first-generation US contrast agent Albunex began to emit acoustic signatures characteristic of nonlinear oscillations at acoustic pressures as low as 0.005 – 0.010 MPa {Krishna & Newhouse 1997}. At even modest Pa (*e.g*., 0.6 MPa at 1 MHz), some contrast agent bubbles can expand to 10 times or larger than their equilibrium radius; under such conditions, inertial collapse is expected. Fragmentation of the inertially-collapsing bubble occurs at some instant near the time of minimum bubble radius {Postema & Schmitz 2007}. Inertial bubble collapse is so rapid and the gas so compressed that the maximum temperature inside the bubble can reach more than 5000 K {Apfel & Holland 1991}. Light emissions {Matula 2003} and reactive free radicals may be produced {Okada *et al.* 2009}. MB rebound and fragmentation can re-emit acoustic energy at many times the excitation pressures. These emissions also contain higher frequency spectral components which can produce local heating. In the present context, however, it appears that the most important determinant of UTMD gene delivery methods is that *bubble expansion and collapse can produce local tissue distortions or damage which are presumed to be the principal mechanism by which* 

*microvascular bioeffects become manifest, and target cell permeabilization occurs.*

of bubble translation {Chahine 1977; Robinson *et al.* 2001}.

would seem rare.

Cavitation bubbles can interact with nearby boundaries; these interactions can produce high velocity fluid jets and induce bubble translation, with the directions of fluid jets and translation depending on the boundary's properties. For rigid boundaries, bubbles closer to the boundary than about twice the fully-expanded radius of the bubble {Kodama & Tomita 2000} may collapse asymmetrically, with a high velocity water jet 'punching through' the bubble and impinging on the rigid boundary. Near rigid, planar boundaries, an oscillating MB translates toward the boundary and if collapse jetting occurs, the jet is directed toward the rigid surface {Plesset & Chapman 1971} with sufficient water hammer pressure to damage the surface {Blake & Gibson 1987}. *In vivo*, however, this condition

MBs driven to oscillate near pressure release boundaries (an air-water interface, for example), collapse jets which occur are directed away from the boundary, as is the direction

Oscillating MBs collapsing near a 'soft', planar boundary (*e.g*., gels or large vessel walls) also undergo translation and jetting, but the directions of jetting and translation can be either toward or away from the boundary, being determined by bubble size, bubble distance from

agents currently under development for imaging or therapy are smaller still, and are considered to be nanobubbles {Krupka *et al.* 2009}. Multi-layered structures in which a gas body stabilized by a lipid monolayer is contained in an aqueous compartment bounded by a lipid bilayer (echogenic liposomes) are also under development. *When used for imaging, US contrast agents are unique amongst contrast agents in that they respond dynamically to the signal used to interrogate tissues for their presence.* It is this same property that is exploited for UTMD therapies or therapy models.

#### **2.3 Microbubble destruction: Shell disruption and shell/gas body fragmentation**

There are three principal mechanisms by which US or time can destroy a shelled MB {Chomas *et al.* 2001}. In increasing order of 'violence' to the MBs, these are: (1) static diffusion, in which gas dissolves into the surrounding host fluid. This may be rapid if the bubble is 'free' (without stabilizing shell) or very slow if a stabilizing shell is present; (2) acoustically-forced shell disruption, which leads to accelerated diffusive loss of gas relative to unperturbed shelled bubbles; and (3) shell fragmentation and rapid loss of gas.

A micron-sized free air bubble can be expected to dissolve in water in ~30 ms {Sarkar *et al.* 2009}, which is why relatively insoluble gases such as perfluorocarbons are now used in modern contrast agents, and a stabilizing shell is employed. The principal mechanism by which bubble shell materials stabilize MBs against diffusion is by reducing the surface tension at the bubble surface. Shell properties also affect the dynamic responses and stability of MBs at Pa below the shell fragmentation threshold {Emmer *et al.* 2007; Ferrara *et al.* 2009}.

Acoustically-forced MB dissolution can occur when the Pa is sufficient to drive MB oscillations which stretch and compress the shell sufficiently to produce relatively small defects in the shell, which may re-seal {Huang 2008}. Very high physical stresses develop in the shells as the MB oscillates, even when the driving pressures are only a few hundred kPa {Stride & Saffari 2003}. Small shell defects lead to more rapid dissolution of the gas body than would otherwise occur, but complete gas dissolution may require many acoustic cycles {Smith *et al.* 2010}. For examples, Optison MBs exposed to 3.5 MHz US of Pa > 0.15 MPa undergo accelerated loss of gas, indicating shell compromise {Porter *et al.* 2006}. Shell disruption Pa thresholds for albumin-stabilized Optison and surfactant stabilized Sonazoid are similar; at 1.1 MHz, these are 0.13 MPa or 0.15 MPa, respectively. At 3.5 MHz, these thresholds are somewhat higher (Optison: 0.48 MPa; Sonazoid: 0.58 MPa) {Chen *et al.* 2003}. Others have reported a somewhat lower threshold (0.15 MPa) for Optison shell disruption at 3.5 MHz {Porter *et al.* 2006}. Likewise, {Borden *et al.* 2005} studied the behavior of lipid monolayer-encapsulated MBs at 2.25 MHz using single-cycle pulses, and found that the Pr threshold for slow MB dissolution to be in the range of 0.4 – 0.6 MPa, and the Pr threshold for MB fragmentation to be 0.8 MPa. A Pa of around 0.3 MPa was also found to be the threshold for soft-shelled contrast agent MB disruption {de Jong *et al.* 2009}. Echogenic liposomes also appear to have two pressure thresholds; one associated with compromise of the shell of the interior gas body, and a higher one associated with disruption of the outer lipid bilayer {Smith *et al.* 2010}. Rapid fragmentation occurs on a time scale of microseconds when Pa is sufficient to buckle and completely rupture the shell {Marmottant *et al.* 2005}, creating free gas bodies which may then dissolve, grow by rectified diffusion, coalesce into larger bubbles, or dissolve, depending stochastically on exposure conditions.

agents currently under development for imaging or therapy are smaller still, and are considered to be nanobubbles {Krupka *et al.* 2009}. Multi-layered structures in which a gas body stabilized by a lipid monolayer is contained in an aqueous compartment bounded by a lipid bilayer (echogenic liposomes) are also under development. *When used for imaging, US contrast agents are unique amongst contrast agents in that they respond dynamically to the signal used to interrogate tissues for their presence.* It is this same property that is exploited for UTMD

**2.3 Microbubble destruction: Shell disruption and shell/gas body fragmentation**  There are three principal mechanisms by which US or time can destroy a shelled MB {Chomas *et al.* 2001}. In increasing order of 'violence' to the MBs, these are: (1) static diffusion, in which gas dissolves into the surrounding host fluid. This may be rapid if the bubble is 'free' (without stabilizing shell) or very slow if a stabilizing shell is present; (2) acoustically-forced shell disruption, which leads to accelerated diffusive loss of gas relative

to unperturbed shelled bubbles; and (3) shell fragmentation and rapid loss of gas.

A micron-sized free air bubble can be expected to dissolve in water in ~30 ms {Sarkar *et al.* 2009}, which is why relatively insoluble gases such as perfluorocarbons are now used in modern contrast agents, and a stabilizing shell is employed. The principal mechanism by which bubble shell materials stabilize MBs against diffusion is by reducing the surface tension at the bubble surface. Shell properties also affect the dynamic responses and stability of MBs at Pa below the shell fragmentation threshold {Emmer *et al.* 2007;

Acoustically-forced MB dissolution can occur when the Pa is sufficient to drive MB oscillations which stretch and compress the shell sufficiently to produce relatively small defects in the shell, which may re-seal {Huang 2008}. Very high physical stresses develop in the shells as the MB oscillates, even when the driving pressures are only a few hundred kPa {Stride & Saffari 2003}. Small shell defects lead to more rapid dissolution of the gas body than would otherwise occur, but complete gas dissolution may require many acoustic cycles {Smith *et al.* 2010}. For examples, Optison MBs exposed to 3.5 MHz US of Pa > 0.15 MPa undergo accelerated loss of gas, indicating shell compromise {Porter *et al.* 2006}. Shell disruption Pa thresholds for albumin-stabilized Optison and surfactant stabilized Sonazoid are similar; at 1.1 MHz, these are 0.13 MPa or 0.15 MPa, respectively. At 3.5 MHz, these thresholds are somewhat higher (Optison: 0.48 MPa; Sonazoid: 0.58 MPa) {Chen *et al.* 2003}. Others have reported a somewhat lower threshold (0.15 MPa) for Optison shell disruption at 3.5 MHz {Porter *et al.* 2006}. Likewise, {Borden *et al.* 2005} studied the behavior of lipid monolayer-encapsulated MBs at 2.25 MHz using single-cycle pulses, and found that the Pr threshold for slow MB dissolution to be in the range of 0.4 – 0.6 MPa, and the Pr threshold for MB fragmentation to be 0.8 MPa. A Pa of around 0.3 MPa was also found to be the threshold for soft-shelled contrast agent MB disruption {de Jong *et al.* 2009}. Echogenic liposomes also appear to have two pressure thresholds; one associated with compromise of the shell of the interior gas body, and a higher one associated with disruption of the outer lipid bilayer {Smith *et al.* 2010}. Rapid fragmentation occurs on a time scale of microseconds when Pa is sufficient to buckle and completely rupture the shell {Marmottant *et al.* 2005}, creating free gas bodies which may then dissolve, grow by rectified diffusion, coalesce into larger bubbles, or dissolve,

therapies or therapy models.

Ferrara *et al.* 2009}.

depending stochastically on exposure conditions.

#### **2.4 Microbubble dynamics: An overview of how gas bodies respond to ultrasound**

At very low Pa, MB volume oscillations are related linearly to Pa. At modest Pas of a few hundred kPa (or less), MB volume oscillations become non-linearly related to Pa, with bubble expansion being relatively slow and bubble collapse much faster, being governed by the inertia of the in-rushing surrounding fluid to a greater extent than by the compressive phase of the applied pressure field; hence the phenomenon of bubbles undergoing acoustically-driven expansions followed by rapid, inertially-dominated collapse is termed 'inertial cavitation'. This typically occurs when a MB has been driven by the rarefaction phase of the acoustic wave to a diameter roughly 2 – 3 times the initial diameter {Chomas *et al.* 2001}. At very modest Pas, inertial cavitation can be stable and sustained for an almost indefinite number of acoustic cycles. For example, {Church & Carstensen 2001} found that surfactant-coated Sonazoid MBs could undergo repetitive inertial collapses and rebounds when driven by 2.5 MHz US at acoustic pressures greater than about 0.3 – 0.4 MPa; for this agent, irreversible post-collapse fragmentation occurred at a Pa of ~ 1.5 MPa. Others have reported that a 3 µm bubble exposed to 2.25 MHz US can be expected to undergo expansion, inertial collapse, and fragmentation at a Pa just over 0.3 MPa {Chomas *et al.* 2001}. The albumin-shelled, first-generation US contrast agent Albunex began to emit acoustic signatures characteristic of nonlinear oscillations at acoustic pressures as low as 0.005 – 0.010 MPa {Krishna & Newhouse 1997}. At even modest Pa (*e.g*., 0.6 MPa at 1 MHz), some contrast agent bubbles can expand to 10 times or larger than their equilibrium radius; under such conditions, inertial collapse is expected. Fragmentation of the inertially-collapsing bubble occurs at some instant near the time of minimum bubble radius {Postema & Schmitz 2007}. Inertial bubble collapse is so rapid and the gas so compressed that the maximum temperature inside the bubble can reach more than 5000 K {Apfel & Holland 1991}. Light emissions {Matula 2003} and reactive free radicals may be produced {Okada *et al.* 2009}. MB rebound and fragmentation can re-emit acoustic energy at many times the excitation pressures. These emissions also contain higher frequency spectral components which can produce local heating. In the present context, however, it appears that the most important determinant of UTMD gene delivery methods is that *bubble expansion and collapse can produce local tissue distortions or damage which are presumed to be the principal mechanism by which microvascular bioeffects become manifest, and target cell permeabilization occurs.*

Cavitation bubbles can interact with nearby boundaries; these interactions can produce high velocity fluid jets and induce bubble translation, with the directions of fluid jets and translation depending on the boundary's properties. For rigid boundaries, bubbles closer to the boundary than about twice the fully-expanded radius of the bubble {Kodama & Tomita 2000} may collapse asymmetrically, with a high velocity water jet 'punching through' the bubble and impinging on the rigid boundary. Near rigid, planar boundaries, an oscillating MB translates toward the boundary and if collapse jetting occurs, the jet is directed toward the rigid surface {Plesset & Chapman 1971} with sufficient water hammer pressure to damage the surface {Blake & Gibson 1987}. *In vivo*, however, this condition would seem rare.

MBs driven to oscillate near pressure release boundaries (an air-water interface, for example), collapse jets which occur are directed away from the boundary, as is the direction of bubble translation {Chahine 1977; Robinson *et al.* 2001}.

Oscillating MBs collapsing near a 'soft', planar boundary (*e.g*., gels or large vessel walls) also undergo translation and jetting, but the directions of jetting and translation can be either toward or away from the boundary, being determined by bubble size, bubble distance from

Ultrasound-Mediated Gene Delivery 219

Optison and Sonazoid were equivalent in effectiveness, producing about four times greater

Bubbles are buoyant, fragile, can coalesce into larger bubbles, or 'disappear' as gas is lost to diffusion. In many *in vivo* studies of UTMD, the MBs and gene vectors are infused slowly; this prolongs the time window in which acoustic treatments can occur, but aggravates the problem of time-dependent changes in MB distribution and delivery {Kaya *et al.* 2009}. It is worth noting that MBs can be destroyed by static pressure or tension; an over-pressure of only ~0.3 MPa can destroy them {Stringham *et al.* 2009}. MBs can also be destroyed by drawing up or injecting an MB suspension *too rapidly* when using a small gauge needle {Talu *et al.* 2008}; thus one can unwittingly alter experimental outcome by 'over-enthusiasm'.

In order to achieve efficient gene expression following ultrasound-mediated gene delivery, multiple barriers need to be overcome to allow pDNA to enter into the nucleus of target cells including penetrating vascular and cellular membranes as well as trafficking through

US contrast agents are intended to be intravascular agents. Whether the gene vector is administered as a simple mixture with the gas body suspension, or is in some way linked to the gas bodies, they are also intravascular agents in vascular UTMD methods. The first barrier encountered is the vascular endothelium. The next are other vascular anatomical features (*e.g.*, the basement membrane, smooth muscle layer, *etc.*) and then the outer cell membrane of the cells one hopes to target. Some intracellular membranous compartments must also be traversed. However, MBs have been used with intramuscular or intraparenchymal injections of vectors, and some successes reported. In UTMD-based gene transfer methods using vascular approaches, to mediate gene therapy *via* acoustic excitation,

Most of the available evidence from *in vivo* studies indicate that vessel permeabilization effects occur principally in the microcirculation; larger vessels are too robust to be penetrated by cavitation events, even if their vascular endothelium can be effectively destroyed by intraluminal inertial cavitation {Hwang *et al.* 2005}. Extravasation of dyes, nanoparticles or macromolecules through microvessels is almost always accompanied by extravasation of red cells (see below); since these have diameters on the order of 6 µm, breaches in the endothelial wall can be quite large. However, there is also some evidence that more subtle effects, such as partial opening of the tight junctions between endothelial cells, can also contribute. Assuming that the gene vector escapes the intravascular compartment and enters the interstitium, it must then enter the surrounding cells; thus the plasma membrane is the second major barrier encountered by the vectors. Lethal effects of cavitation occurring in the cardiac microcirculation can extend outward into the myocardium {Miller *et al.* 2011}; there is good reason to expect that sub lethal poration of cells located within a few cell diameters of the intravascular cavitation event(s) also occurs.

Evans blue [EB] is an azo dye which binds serum albumin with high affinity, and is normally unable to pass through the endothelium. Extravasation of EB through reversible or

**3. Breaching the physical barriers to gene delivery using ultrasound** 

gas bodies must first exert their influence from within the vascular lumen.

**3.1 Extravasation of dyes, nanoparticles and cells** 

expression than did SonoVue {Alter *et al.* 2009}.

different intracellular compartments.

**2.6 Sensitivity of stabilized gas bodies to technique** 

the boundary, and the mechanical properties of the boundary (see. *e.g*., {Kodama & Tomita 2000; Shima *et al.* 1989}. Deformation of the elastic boundary stores energy which is 'returned' as the bubble collapses; this creates a hydrodynamic pressure gradient which in turn produces fluid flow away from the elastic boundary. If the pressure gradient is large enough, bubble collapse jets directed away from the boundary, and bubble translation away from the boundary, can result {Blake & Gibson 1987}. Liquid jets were observed in an early experimental study of acoustically forced MB behavior in a 200 µm diameter cellulose tube {Postema *et al.* 2004}. In vessel-simulating gel tunnels, at 1.7 MHz, the Pa at which Optison MBs began to emit broadband noise characteristic of inertial collapse and rebound was weakly dependent on tunnel diameter (~0.8 MPa in 90 µm diameter tunnels *vs*. ~0.6 MPa in 800 µm tunnels) {Sassaroli & Hynynen 2007}. Single-MB dynamics observations in actual microvessels were acquired using a rat cecum model; MBs translated toward the vessel walls and a toroidal bubble morphology consistent with the formation of a microjet was observed {Caskey *et al.* 2007}.

Where bioeffects attributable to inertial cavitation occur, cavitation jets directed toward cells or tissues have long been assumed to be causative. Vessel 'stretching' during bubble expansion has also been proposed as a mechanism of vessel damage {Miao *et al.* 2008}. Very high speed imaging of MBs driven to growth and inertial collapse in microvessels of rat mesentery {Chen *et al.* 2010; Chen *et al.* 2011} indicate that (1) MB collapse jets form frequently in the intravascular environment, (2) jets are typically directed *away* from the nearest blood vessel wall, (3) blood vessel expansion in response to bubble growth is minimal, and (4) tissues 'follow' the collapsing bubbles such that inward vessel distortions are much greater than outward distortions, with inward vessel wall motion having speeds of 5 – 10 m/s and the events occurring on microsecond time scales. This is perhaps one mechanism by which vessel permeabilization occurs. In any case, permeabilization of biological transport barriers is associated with some cell killing, both *in vitro* {Brayman *et al.* 1999} and *in vivo* {Ferrara 2008; Miller *et al.* 2011; Price *et al.* 1998; Skyba *et al.* 1998}

Finally, it is worth mentioning that MB-excitation techniques for cellular permeabilization need not use US to achieve MB activation or cell permeabilization; targeted laser illumination can also be used effectively. For a recent research paper on this topic, using a liposome MB contrast agent, see {Zhou *et al.* 2010}.

#### **2.5 Effect of different shell compositions or agents on gene delivery**

It is difficult to compare the efficacy of different contrast agents in US-mediated gene delivery between different studies, because MB type, concentration as injected, rate of injection, total dose, animal model, US exposure source, exposure conditions, and even the way the MBs are handled vary widely and can influence outcome. MB concentration is a determinant of UTMD-mediated pDNA expression in mouse liver {Miao *et al.* 2005; Shen *et al.* 2008}. Different agents have different 'native' concentrations, and agents with similar gas content but dissimilar shells can have different efficacies in enhancing gene delivery under otherwise comparable conditions (*cf*. effects of Optison and PESDA MBs; {Pislaru *et al.* 2003}). Relatively large differences between contrast agent types in UTMD gene delivery have been observed, but the reasons are not always clear. With equalized MB concentrations and either Optison, SonoVue or Sonazoid contrast agents, UTMD-mediated transfer of phosphorodiamidate morpholino oligomer to a dystrophin-deficient murine heart model induced dystrophin-positive cells in the hearts when harvested 7 days after treatment.

the boundary, and the mechanical properties of the boundary (see. *e.g*., {Kodama & Tomita 2000; Shima *et al.* 1989}. Deformation of the elastic boundary stores energy which is 'returned' as the bubble collapses; this creates a hydrodynamic pressure gradient which in turn produces fluid flow away from the elastic boundary. If the pressure gradient is large enough, bubble collapse jets directed away from the boundary, and bubble translation away from the boundary, can result {Blake & Gibson 1987}. Liquid jets were observed in an early experimental study of acoustically forced MB behavior in a 200 µm diameter cellulose tube {Postema *et al.* 2004}. In vessel-simulating gel tunnels, at 1.7 MHz, the Pa at which Optison MBs began to emit broadband noise characteristic of inertial collapse and rebound was weakly dependent on tunnel diameter (~0.8 MPa in 90 µm diameter tunnels *vs*. ~0.6 MPa in 800 µm tunnels) {Sassaroli & Hynynen 2007}. Single-MB dynamics observations in actual microvessels were acquired using a rat cecum model; MBs translated toward the vessel walls and a toroidal bubble morphology consistent with the formation of a microjet was

Where bioeffects attributable to inertial cavitation occur, cavitation jets directed toward cells or tissues have long been assumed to be causative. Vessel 'stretching' during bubble expansion has also been proposed as a mechanism of vessel damage {Miao *et al.* 2008}. Very high speed imaging of MBs driven to growth and inertial collapse in microvessels of rat mesentery {Chen *et al.* 2010; Chen *et al.* 2011} indicate that (1) MB collapse jets form frequently in the intravascular environment, (2) jets are typically directed *away* from the nearest blood vessel wall, (3) blood vessel expansion in response to bubble growth is minimal, and (4) tissues 'follow' the collapsing bubbles such that inward vessel distortions are much greater than outward distortions, with inward vessel wall motion having speeds of 5 – 10 m/s and the events occurring on microsecond time scales. This is perhaps one mechanism by which vessel permeabilization occurs. In any case, permeabilization of biological transport barriers is associated with some cell killing, both *in vitro* {Brayman *et al.*

1999} and *in vivo* {Ferrara 2008; Miller *et al.* 2011; Price *et al.* 1998; Skyba *et al.* 1998}

**2.5 Effect of different shell compositions or agents on gene delivery** 

liposome MB contrast agent, see {Zhou *et al.* 2010}.

Finally, it is worth mentioning that MB-excitation techniques for cellular permeabilization need not use US to achieve MB activation or cell permeabilization; targeted laser illumination can also be used effectively. For a recent research paper on this topic, using a

It is difficult to compare the efficacy of different contrast agents in US-mediated gene delivery between different studies, because MB type, concentration as injected, rate of injection, total dose, animal model, US exposure source, exposure conditions, and even the way the MBs are handled vary widely and can influence outcome. MB concentration is a determinant of UTMD-mediated pDNA expression in mouse liver {Miao *et al.* 2005; Shen *et al.* 2008}. Different agents have different 'native' concentrations, and agents with similar gas content but dissimilar shells can have different efficacies in enhancing gene delivery under otherwise comparable conditions (*cf*. effects of Optison and PESDA MBs; {Pislaru *et al.* 2003}). Relatively large differences between contrast agent types in UTMD gene delivery have been observed, but the reasons are not always clear. With equalized MB concentrations and either Optison, SonoVue or Sonazoid contrast agents, UTMD-mediated transfer of phosphorodiamidate morpholino oligomer to a dystrophin-deficient murine heart model induced dystrophin-positive cells in the hearts when harvested 7 days after treatment.

observed {Caskey *et al.* 2007}.

Optison and Sonazoid were equivalent in effectiveness, producing about four times greater expression than did SonoVue {Alter *et al.* 2009}.

#### **2.6 Sensitivity of stabilized gas bodies to technique**

Bubbles are buoyant, fragile, can coalesce into larger bubbles, or 'disappear' as gas is lost to diffusion. In many *in vivo* studies of UTMD, the MBs and gene vectors are infused slowly; this prolongs the time window in which acoustic treatments can occur, but aggravates the problem of time-dependent changes in MB distribution and delivery {Kaya *et al.* 2009}. It is worth noting that MBs can be destroyed by static pressure or tension; an over-pressure of only ~0.3 MPa can destroy them {Stringham *et al.* 2009}. MBs can also be destroyed by drawing up or injecting an MB suspension *too rapidly* when using a small gauge needle {Talu *et al.* 2008}; thus one can unwittingly alter experimental outcome by 'over-enthusiasm'.
