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Ultrasound MicrobubbleContrast Agents:Fundamentals andApplication to Gene andDrug DeliveryKatherine Ferrara,1 Rachel Pollard,2

and Mark Borden1

1Department of Biomedical Engineering and 2Department of Surgical andRadiological Sciences, University of California, Davis, California 95616-8686;email: kwferrara@ucdavis.edu

Annu. Rev. Biomed. Eng. 2007. 9:415–47

The Annual Review of Biomedical Engineering isonline at bioeng.annualreviews.org

This article’s doi:10.1146/annurev.bioeng.8.061505.095852

Copyright c© 2007 by Annual Reviews.All rights reserved

1523-9829/07/0815-0415$20.00

Key Words

controlled release, DNA, chemotherapy, cavitation, radiationforce, sonoporation

AbstractThis review offers a critical analysis of the state of the art of med-ical microbubbles and their application in therapeutic delivery andmonitoring. When driven by an ultrasonic pulse, these small gasbubbles oscillate with a wall velocity on the order of tens to hun-dreds of meters per second and can be deflected to a vessel wallor fragmented into particles on the order of nanometers. Whilesingle-session molecular imaging of multiple targets is difficult withaffinity-based strategies employed in some other imaging modali-ties, microbubble fragmentation facilitates such studies. Similarly, afocused ultrasound beam can be used to disrupt delivery vehicles andblood vessel walls, offering the opportunity to locally deliver a drugor gene. Clinical translation of these vehicles will require that cur-rent challenges be overcome, where these challenges include rapidclearance and low payload. The technology, early successes with drugand gene delivery, and potential clinical applications are reviewed.

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Contents

INTRODUCTION. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 416PHYSICOCHEMISTRY OF MICROBUBBLES . . . . . . . . . . . . . . . . . . . . . . 417

Microbubble Stability . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 417Properties of the Lipid Shell . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 419Other Shell Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 420Targeting . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 421Types of Ligands . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 423Payload . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 423

MICROBUBBLE PHYSICS DURING INSONIFICATION. . . . . . . . . . . 424Microbubble Oscillation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 424Radiation Force . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 426

MICROBUBBLE-SPECIFIC IMAGING TECHNIQUES. . . . . . . . . . . . . 427IMAGING THE RESPONSE TO THERAPY . . . . . . . . . . . . . . . . . . . . . . . . 429CHANGES IN CELL MEMBRANE AND VASCULAR

PERMEABILITY . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 430APPLICATIONS: POTENTIAL IMAGING AND

THERAPEUTIC TARGETS . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 431Thrombolytic Therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 433Anti-Cancer Therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 433Atherosclerotic Therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 434Gene Therapy. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 435

CHALLENGES AND LIMITATIONS. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 436

INTRODUCTION

Small gas bubbles, used to enhance ultrasound contrast, were first reported byGramiak & Shah (1). Air bubbles introduced without a stabilizing shell were veryshort lived, and therefore methods to stabilize the gas-liquid interface were devel-oped. Simultaneously, low diffusivity gases were introduced to further increase themicrobubble circulation time. Clinical ultrasound systems now incorporate contrast-specific imaging modes designed to take advantage of unique nonlinear propertiesof these small gas nuclei. Today, these stabilized gas bubbles are used to enhancethe reflectivity of perfused tissues in applications spanning cardiology and radiology.Although molecularly targeted agents have not yet been applied clinically, success-ful preclinical studies have demonstrated application in angiogenesis and vascularinflammation, and are described in this review.

In the absence of an exogenous contrast agent, the safety of ultrasound energy hasbeen widely studied, resulting in guidelines designed to control thermal and mechan-ical biological effects (2). Given that microbubble contrast agents can be expanded,moved, or fragmented by ultrasound, insonation of these agents requires additionalsafety consideration. Alterations in cell membrane and vascular permeability resulting

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from in vitro and in vivo insonation of contrast agents are now well established (3–5).These studies present cautionary limitations for imaging studies and new opportuni-ties for drug delivery. Creation of a clinically optimized microbubble-based drug orgene delivery vehicle is challenging because there exists a vast parameter space of po-tential shell materials, targeting ligands, cargos, and ultrasound parameters. Herein,engineering analyses of both the microbubble shell and gas core are reviewed. As-pects of the physicochemical properties and ultrasound response of microbubbles arepresented in the first two sections. Later sections focus on their imaging and potentialtherapeutic applications. Finally, we conclude with a section on current limitationsand challenges of ultrasound-guided gene and drug delivery with microbubbles.

PHYSICOCHEMISTRY OF MICROBUBBLES

Microbubble Stability

A newly formed microbubble without encapsulation will dissolve spontaneouslyand nearly instantaneously as a consequence of surface tension at the gas-liquidinterface. Surface tension (σ ) is a fundamental property of a fluid interface betweentwo immiscible phases, and when the interface is curved, results in a higher pressureon the concave side. The pressure drop across a bubble interface (�P) is given bythe Laplace equation (6):

�P = Pb − Pa = 2σ

r, (1.1)

where Pb is the pressure inside the bubble, Pa is the hydrostatic pressure outside thebubble and r is the bubble radius. The high curvature of a microbubble renders a sig-nificant pressure drop, on the order of 1 bar for a 2-μm-diameter bubble, which drivesgas into the surrounding medium. An encapsulating shell—a barrier between theaqueous and gas phases—is necessary to sustain the gas cavity. The shell contributestwo stabilizing components: a resistance to gas leaving the core and a reduction insurface tension, as modeled in a modified Epstein-Plesset (EP) equation (7, 8),

−drdt

= Lr/Dw + Rshell

(1 + 2σshell/Par − f

1 + 3σshell/4Par

), (1.2)

where L is Ostwald’s coefficient, Dw is the gas diffusivity in water, Rshell is theresistance of the shell to gas permeation, σ shell is the surface tension of the shell, andf is the ratio of the gas concentration in the bulk medium versus that at saturation.This model assumes a perfectly spherical geometry throughout dissolution andneglects changes in the shell properties as it deforms to accommodate the shrinkinggas core. However, it provides useful simulations to illustrate the shell’s stabilizingeffects.

The lifetime of a free microbubble is simulated by neglecting the shell resis-tance and applying the surface tension for a clean gas-liquid interface. As shown inFigure 1a, a free air microbubble completely dissolves in less than a second in a gas-saturated liquid ( f = 1). One approach to increase stability has been to use insolublegases, such as perfluorocarbons, which have water permeation resistances (L−1D−1

w )

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Air PFB Air PFB

0.5

1.0

1.5

2.0

2.5

3.0

010-3 10-2 10-1 100 101 102 10-2 10-1 100 101 104103102 105 106

f = 1

a Free microbubble dissolution b Lipid-coated microbubble dissolution

0.5

1.0

1.5

2.0

2.5

3.0

0

f = 0

Time (s)

Bu

bb

le r

adiu

s (µ

m)

Bu

bb

le r

adiu

s (µ

m)

Time (s)

Figure 1Calculated microbubble dissolution kinetics based on the modified-EP equation. (a)Radius-time curves of a free microbubble composed of air or perfluorobutane (PFB). Modelparameters were σshell = 72 mN m−1, Rshell = 0, Pa = 101.3 kPa, and f = 1 (i.e., saturation).Diffusion parameters for air were L = 0.02 and Dw = 2 × 10−5 cm2 s−1; those for PFBwere L = 0.0002, Dw = 0.7 × 10−5 cm2 s−1. (b) Radius-time curves of lipid-coatedmicrobubbles in degassed water ( f = 0). Model parameters were the same as above, exceptRshell = 104 s m−1 for air and 107 s m−1 for PFB and σshell = 0 mN m−1. All curves weregenerated in MATLAB using a standard fourth-order Runga-Kutta algorithm.

that are several orders of magnitude higher than air (9, 10). For example, the waterpermeation resistance of perfluorobutane (PFB, n-C4F10) is ∼300 times greater thanthat of air. Figure 1a shows, however, that the lifetime of a free PFB microbubble isless than a minute, which is far too short for use as a medical device.

For a suitable shelf-life (days to months), the microbubble shell must be solidto eliminate surface tension and impart a significant permeation resistance. Currentultrasound contrast agents typically comprise solid shells of various compositions,as discussed below. Removing the Laplace-overpressure term eliminates the impetusfor dissolution in saturated media, as is easily demonstrated by Equation 1.2. Thisallows long-term storage of microbubbles, which are stable indefinitely in a sealedvial.

The effect of shell permeation resistance is enhanced for higher molecular weightgases. Permeation through surfactant monolayer films occurs by a thermally acti-vated transition state (11–13), rather than partition-diffusion. The microbubble shellresistance for a high-molecular-weight gas, such as PFB, can be estimated from theoxygen resistance according to the relation (11, 14)

RPFBshell

Roxygenshell

= exp(

π�

4kT(a2

PFB − a2oxygen)

), (1.3)

where a is the collision diameter of the permeating species, � is the surface pressure(� = σ − σshell), k is Boltzmann’s constant, and T is the absolute temperature. The

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values of a for oxygen and n-C4F10 are 0.36 and 0.81 nm, respectively (15). Usingthese values, the ratio of resistances for PFB to air is estimated to be ∼1400. Thisis an important point because it often is taken for granted that the shell resistanceis negligible (9, 10, 16). Indeed, for high-molecular-weight gases, shell resistance isan important stabilizing term. For the sake of illustrating this effect, a comparisonis made using the modified-EP equation under degassed conditions ( f = 0). Theoxygen permeability of solid lipid monolayers was measured to be approximately 102

s/cm (12, 17), and this value is used with Equation 1.3 to estimate the value for PFBat 105 s/cm. Figure 1b shows that the estimated microbubble lifetime is significantlyenhanced, by five orders of magnitude, by the shell’s diffusion impedance to PFB.The low solubility and slow dissolution kinetics of lipid-coated, perfluorocarbon-filled microbubbles allows for the high degree of dilution necessary to modify themicrobubbles for targeting and drug delivery.

Properties of the Lipid Shell

Lipids are a large class of compounds that consist of one or more hydrocarbon orfluorocarbon chains covalently attached to a hydrophilic headgroup, usually by aglycerol backbone. Lipid shells are less stable than other shell types (e.g., polymers),but they are easier to form and yield a more echogenic microbubble.

Lipids spontaneously adsorb from soluble aggregates (i.e., micelles and vesicles)to the gas-liquid interface and self assemble into a monolayer coating (18, 19). Atthe nanoscale, the molecules are oriented such that the hydrophobic tails face the gasphase and interact via hydrophobic and dispersion forces, which can be modulatedby increasing or decreasing chain length. Lipids that are below the main phase tran-sition temperature form highly condensed shells. Increasing chain length was foundto diminish surface tension (16) and increase surface viscosity (20), gas permeationresistance (8, 12, 17), and buckling stability (8, 21, 22) of the shell, resulting in morerobust microbubbles.

Recent findings have changed the prevailing paradigm concerning the structureof the lipid shell; it is now realized to be a complex, multiphase structure (Figure 2).Pioneering work by Kim et al. showed that the shell consists of planar microdomains(grains) separated by defects (grain boundaries), which influence the mechanical prop-erties (20, 23). Pu et al. later showed that defect density also affects the gas permeability(17).

Research by Borden et al. also showed that the grain boundary regions are aseparate, more labile phase that is enriched in certain monolayer components, such aslipopolymers, whereas the microdomains are composed mainly of lecithin (24). Bothphases are necessary to stabilize the microbubbles (21, 25). Because the monolayeris compressed beyond the equilibrium spreading pressure, meta-stability of the shellis defined by the collapse rate (i.e., the speed of the 2-D-to-3-D transition). Thetwo-phase system shifts the collapse transition from vesiculation to the more stablemode of buckling and folding (26). The folds, or wrinkles, can be stably attachedand continuous to the monolayer shell (8, 21), and external shear forces can act torespread the shell over the elongated bubble surface (23, 27).

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b c

d

a

e f

Figure 2Phase separation in lipid-coated microbubbles. Shown are individual microbubbles imagedwith epifluorescence or confocal microscopy that show preferential partitioning to differentregions of the shell. Different species were labeled with dye, including (a) DiIC18 membraneprobe, (b) DiOC18 membrane probe, (c) FITC-neutravidin bound to DSPE-PEG2000-biotin,(d ) avidin-coated fluorescent nanobeads bound to DSPE-PEG2000-biotin,(e) NBD-cholesterol, and ( f ) YoYo-1-labeled DNA adsorbed to a cationic lipid.

Lipids can be chosen from a virtually unlimited library to present a shieldingbrush layer, express ligands, and bind to drugs, genes and other compounds. Thelipopolymer and lecithin species can be mixed with ligand-conjugated or chargedlipids, for example, and the surface microstructure and brush architecture can beengineered to yield novel properties. Additionally, lipid shells are highly compliantto severe area dilation and compression during ultrasound-induced oscillation. Thelipid shell readily expands, ruptures, reseals, compresses, buckles, and respreads witheach acoustic cycle (22, 28–32).

Other Shell Materials

In the 1980s, ultrasound contrast agents were coated with an adsorbed layer ofsaccharide (33) or protein (34). The albumin-coated microbubbles Albunex® andOptisonTM (GE Healthcare) were the first commercially available, FDA-approvedcontrast agents. More recently, protein-shelled microbubbles have been functional-ized to carry targeting ligands (35) and genetic payloads (36, 37). Albumin shells tendto be rigid (28), however, and introduce the typical immunogenicity issues associatedwith animal-derived materials.

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Several methods have been described to encapsulate gas in a polymer shell.Alignate-shelled microbubbles have been created by dispersion and ionic gelation(38). Several investigators have used organic solvents to dissolve and disperse thepolymer, followed by resuspension of lyophilized products to form hollow polymercapsules (39–41). Partial filming of surfactant-coated microbubbles with nanoparti-cles has been described by Schmidt & Roessling (42). Polymerization at the air-liquidinterface during agitation of an acidic medium was described by Paradossi et al. (43).Each of these methods has produced microbubbles with enhanced stability. However,chain entanglement and covalent bonds inherent in the polymer shells dampen theoscillation of the gas core (32, 44) until the shell is fractured by sufficient expansion.

Targeting

Physicochemical interactions can be used to increase the specificity of targeting. Asimple example is the use of surface charge. Lindner et al. found that cationic mi-crobubbles persist in the microcirculation of tissue undergoing ischemia/reperfusionand inflammation owing to interactions with the innate immune system (45). Elec-trostatic interactions generally are not sufficiently specific, however, given the com-plexity of biological milieu. Ligand-receptor interactions, on the other hand, yieldhigh specificity in biological media. In this case, the microbubble surface is deco-rated with ligands that bind specifically to receptors on cells lining the lumen ofvessels.

As discussed above, a lipopolymer is necessary to form stable microbubbles. Thepresence of the polymer brush necessitates a spacer between the ligand and monolayershell for the ligand to interrogate its receptor on an apposing surface (46). Typically,the ligand is tethered with a spacer that is equal in length or longer than the chains ofthe surrounding brush. This gives maximum exposure of the ligand to the biologicalmilieu.

Surface architectures designed to maximize ligand exposure to the target tissue alsorun the risk of boosting presentation of an immunogenic compound that leads to earlyparticle clearance or, worse, a hypersensitivity response. For example, unpublisheddata in our laboratory clearly show that biotin-conjugated lipopolymer present onmicrobubbles activates the complement system in humans and mice (M. A. Bordenand K. W. Ferrara, unpublished data). More research is necessary to test whethertethered antibody or peptide ligands also elicit an immune response. To account forthe immunogenic effect, Borden et al. (47) showed that the ligand can be concealedby a polymeric overbrush to enhance circulation half-life, and then it can be locallyrevealed for binding to the target by ultrasound radiation force. A discussion ofultrasound radiation force is provided in a subsequent section.

There are two basic methods of attaching ligands to the microbubble surface: ei-ther through a direct covalent bond or through biotin-avidin linking. Biotin-avidinlinkage is a straightforward technique in which a biotinylated ligand is coupled toa biotinylated microbubble via an avidin bridge. Although biotin-avidin linkage isuseful for proof-of-concept and preclinical targeting studies, the immunogenicityprecludes it from translation to humans. Covalent attachment is more desirable and

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Microbubble shell

Receptor surfacea b Receptor surface

Microbubble shell

AvidinAvidin

Biotin

Figure 3Ligand architectures. Cartoons show different schemes for the presentation of various ligandtypes: (a) small hydrophilic ligands can be covalently attached to the distal end of the carrierlipopolymer. The diffusion of the ligand is dictated by the polymer chain dynamics; (b) large,protein ligands can be attached by biotin (red ) and avidin (yellow) linkage. The large size(60 kDa) and multiple binding pockets in the avidin create a sort of scaffold that is supportedby the polymeric brush (2–5 kDa).

can be performed prior or subsequent to creation of the microbubble shell. Strate-gies for coupling onto preformed microbubbles include binding an amino groupof the ligand to a carboxyl group on the microbubble shell by carbodiimide andN-hydroxysulfosuccinimide or, alternatively, binding a thiol group on the ligand to amaleimide on the microbubble shell. More details on the coupling chemistry can befound in a recent review by A.L. Klibanov (48). For lipid-coated agents, the advantageof using a preformed ligand-lipopolymer is that fewer steps are required in the clinicalsetting between microbubble production and administration into the patient. Withpostproduction linkage, however, a more efficient use of ligand is achieved througha series of modifications to preformed microbubbles.

Choice of the appropriate coupling chemistry and sequence of modifications de-pends on the type of ligand. An important consideration is ligand size and its effect onbioavailability (Figure 3). Small, hydrophilic molecules, such as metabolites and pep-tides, can be conjugated directly to the polymer spacer without significantly affectingthe polymer dynamics (49). In contrast, large protein ligands, such as antibodies, aresusceptible to denaturation owing to shear stresses and organic solvents involved withmicrobubble dispersion. For this reason, antibodies (∼120 kDa) typically are conju-gated to the preformed microbubble surface by biotin-avidin linking. The resultingcomplex resembles more of a rigid scaffold than a free polymer chain (50), with theligand well separated from the polymer brush (∼5 kDa) by the bulky avidin molecule(∼60 kDa).

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Types of Ligands

Several classes of ligands have been conjugated onto microbubbles, including an-tibodies, peptides, and vitamins. Monoclonal antibodies, particularly those in theimmunoglobulin-γ (IgG) family, have been used extensively for targeting cell sur-face receptors. Monoclonal antibodies are versatile and have a binding affinity in thenanomolar to picomolar range. However, they tend to be immunogenic when derivedfrom murine origins. Antibody production for targeted imaging and drug deliveryalso tends to be expensive and time consuming with binding activity varying frombatch to batch. Other limitations of antibodies as targeting agents include a limitedshelf-life and temperature sensitivity (51).

Peptides are smaller molecules that offer chemical stability and low immunogenic-ity. The recent development of combinatorial peptide library methods has rapidlyadvanced the use of peptides as targeting ligands. A class of ligands that have not yetbeen used for targeting microbubbles is aptamers. Aptamers are RNA- or DNA-basedligands that bind targets with exceptional affinity and specificity. These ligands arecreated using a process of systematic evolution of ligands by exponential enrichment(SELEX) (51, 52). Because this process is based on chemical synthesis, some of thelimitations encountered with antibody ligands are avoided. A downside is the naturalactivity of RNAses/DNAses.

Payload

In addition to targeted imaging, ultrasound with microbubbles can be used to delivera therapeutic payload. The simultaneous release and enhancement of vascular perme-ability, as discussed in more detail below, are attributes that are unique to ultrasoundmicrobubble technology. A key component to engineering a microbubble formula-tion for intervention is loading the therapeutic agent onto the shell. The gas core isessentially a void that does not sequester organic compounds, and the lipid shell istoo thin (∼3 nm) to hold a sufficient cargo. One approach to enhance loading wasto introduce oil into the lipid shell that dissolves hydrophilic or lipophilic drugs (53).This technique of forming acoustically active lipospheres (AALs) has shown successin delivering chemotherapeutics in vitro (53).

Charged therapeutics, such as DNA or RNA, can be coupled electrostaticallyonto the shell when cationic lipids or denatured proteins are present. This techniquehas been used extensively for gene transfection experiments. The loading capacityobserved experimentally for lipid-coated microbubbles is on the order of 0.01 pg/um2

(34, 55). We are currently testing an approach to improve loading capacity throughthe use of multiple strata to enhance the total surface area for electrostatic coupling,and preliminary data show a greater than tenfold increase in DNA loading capacityfor five paired layers (M.A. Borden & K.W. Ferrara, unpublished data).

Loading on the microbubble surface also can be achieved using ligand-receptorinteractions. For example, Lum et al. recently reported a study in which nanopar-ticles were bound to the shell by biotin-avidin linkage (56). The solid polystyrenenanoparticles served as a model system, which could be replaced by biodegradable

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nanoparticles loaded with drugs or genes. Alternatively, soft nanoparticles, such asliposomes, have been loaded successfully onto microbubbles (48, 57). These re-sults suggest a modular approach for loading, whereby the therapeutic compoundis first loaded into a nanoparticle compartment, which is then loaded onto a mi-crobubble carrier. Such an approach offers a versatile platform that can be tailoredto the hydrophobicity, size, and release requirements of the specific therapeuticagent.

MICROBUBBLE PHYSICS DURING INSONIFICATION

Microbubble Oscillation

During an acoustic pulse, the highly compressible microbubble expands and contractswith the applied pressure rarefaction and compression. Capturing the expansion andbreakup of the agent requires a shutter duration on the order of tens to hundreds ofnanoseconds (Figure 4a). Observation of the microbubble collapse requires a shut-ter duration on the order of picoseconds where a peak wall velocity of hundreds ofmeters per second is observed (29). In the absence of a nearby boundary, sphericaloscillations are observed, and the effect of ultrasound on contrast agents is approxi-mately characterized by measuring the radial oscillations (radius-time curve) of theinsonified agent using a radius-time “streak” image of a single line through the cen-ter of a microbubble (Figure 4b) (29, 30). For lipid-shelled microbubbles driven ata pressure up to several hundred kilopascals, the observed oscillation can be accu-rately predicted by Rayleigh-Plesset analysis (Figure 4c) (29). Although the appliedultrasound pulse is typically sinusoidal, radius-time images of microbubbles driven ata pressure greater than tens of kilopascals demonstrate that the oscillation containsa rich range of frequency components. Furthermore, the mapping of the radial os-cillation into an ultrasound echo is dependent on the radial oscillation and its firstand second time derivatives, accentuating higher frequency components. Thus, thereceived echoes from oscillating microbubbles contain submultiples and multiples ofthe transmission frequency spectrum.

Owing to their size-dependent resonance behavior, microbubbles with a diameteron the order of a few microns differ greatly in their response to ultrasound frequen-cies, which are on the order of megahertz. Parameters such as pulse phase, centerfrequency, and acoustic pressure alter the relative expansion and potential for frag-mentation and can be optimized based on high-speed photography (Figure 5). Inver-sion of the transmitted pulse does not result in simple phase inversion of the radial os-cillation or echo (Figure 5a). Decreasing the transmission center frequency increasesthe maximum diameter achieved during expansion and the likelihood of fragmenta-tion (Figure 5b). Depending on the microbubble diameter and ultrasound frequencyand pressure, the period corresponding to one entire microbubble oscillation may ex-tend beyond the period of the transmitted pulse, resulting in echoes with componentsat submultiples (subharmonics) of the transmitted frequencies (Figure 5c).

Lipid-shelled agents have been studied most frequently and their oscillation canbe predicted (29) and compared with that of albumin-shelled agents (58). Polymer

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Driving waveform

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Figure 4Oscillation of a microbubble in response to an ultrasound field. (a) Sequence oftwo-dimensional images, in which microbubble oscillates in response to a 2.25-MHz centerfrequency pulse, acquired with a camera shutter duration of 10 nanoseconds. (b) Concept of aradius-time image. Microbubble oscillates in response to the driving pressure and adistance-time image is acquired of a single line through the center of the microbubbledemonstrating the increased and decreased diameter. (c) Radius-time image of a microbubbleoscillating in response to a 2.25 MHz pulse with a peak negative pressure of 360 kPa.Transmitted pulse is shown in yellow, and predicted radius-time curve is shown in red.(Figure 4b reprinted with permission from Reference 29).

shells undergo complex changes with insonation and models for their behavior areunder development (59). Models and simulations of contrast agent oscillation haverecently been extended to microbubbles constrained within small blood vessels or nearboundaries (60–64). Within vessels with a diameter of about ten microns, microbubbleoscillation amplitude is decreased, and fragmentation is not observed by a singleultrasound pulse.

Destruction of microbubble contrast agents has been observed during ultrasonicexcitation. The mechanisms of destruction include fragmentation of the bubble intosmaller bubbles and/or dissolution of the encapsulated gas. Relative expansion of a

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0°

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Figure 5Streak images of microbubble oscillation in response to varied pulse parameters. (a) Effect ofthe transmission phase of a single-cycle pulse. As compared with transmission of ararefaction-first (180◦) pulse, when the compressional half-cycle is transmitted first (0◦ pulse),expansion is longer, contraction occurs less rapidly, and the microbubble remains intact.(1) and (3) are streak images, (2) and (4) two-dimensional images acquired after insonation.(b) Effect of the transmission center frequency for matched peak negative pressure. Ascompared to a center frequency of 3.5 MHz, with a transmission center frequency of 1.5MHz, expansion is larger and the microbubble fragments. (1) and (2) are streak images, (3–6)two-dimensional images (3) and (5) acquired at peak expansion, (4) and (6) acquired afterinsonation. (c) Effect of the transmission pressure with a transmission center frequency of2.25 MHz. Streak images demonstrating harmonic and subharmonic oscillation ofmicrobubble-driven with lower and higher transmission pressure, respectively. (Figure 5reprinted with permission from References 30, 154).

microbubble is defined as the maximum radius, Rmax, divided by the resting radius,Ro, and has been correlated with the cavitation threshold and fragmentation (29, 30).

Radiation Force

The effects of radiation force were first reported in 1906 by C.A. Bjerknes and hisson V.K.F. Bjerknes when they observed the attraction and repulsion of air bubbles ina sound field (65). Over the next 80 years, several research groups published theoret-ical analyses of radiation force on spheres, noting both a primary radiation force thattranslates the microbubble along the beam axis and a secondary radiation force re-sulting in microbubble attraction or repulsion (66–69). In these cases, the air bubblesstudied were on the order of tens of microns in radius, and the acoustic frequency wason the order of tens of kilo-Hertz. More recent studies indicated that a strong radi-ation force effect is generated on micron-sized ultrasound contrast agents by clinicalultrasound instruments (53, 56, 60, 70–72). These agents are deflected a distance onthe order of a fraction of a micron with each acoustic cycle when the driving pulse isnear their resonance frequency. Although transmission of a pulse with a high-pressureamplitude would increase the deflection produced by radiation force, microbubblesare fragmented by a pressure on the order of hundreds of kiloPascals. Therefore,

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a very long pulse, with a low pressure and frequency near the bubble’s resonancefrequency is used. Owing to the polydisperse nature of commercial microbubblecontrast agents, high-time-frequency bandwidth pulses have also been considered tomaximize the volume of deflected microbubbles.

MICROBUBBLE-SPECIFIC IMAGING TECHNIQUES

With the addition of microbubble contrast agents, ultrasound becomes sensitive tocapillary-sized vessels and very low flow rates, while maintaining the ability to de-tect morphological information from traditional B-mode imaging. Because they arehighly compressible and result in strong scattering of ultrasound, microbubbles ap-pear bright on an ultrasound image (30). When insonified, the expansion and con-traction of these agents results in the generation of nonlinear signals (30, 73).

Power Doppler imaging involves the transmission and reception of a train of ul-trasound pulses, where the motion of scatterers between pulses is used to detect bloodflow. Combining power Doppler with ultrasound contrast agents results in improveddetection of small vessels (Figure 6). A strong correlation was found between histo-logic microvascular density (MVD) and the number of intratumoral vessels identifiedwith 2- and 3-D power Doppler sonography of human breast masses (74). Anotherstudy used the ratio of enhanced pixels to total pixels in the tumor to follow antiangio-genic treatment of xenografted tumors in mice. The signal-pixel rate was significantlydecreased in the treated versus control group and correlated with MVD (75).

Various other approaches have been described to accentuate the nonlinear contrastagent echoes and repress the echoes generated from surrounding tissue. Harmonicimaging is a broad category of techniques that share the common feature of sendingan incident beam at one frequency and listening for returned echoes at a harmonic(an integer multiple) of the incident beam (Figure 7) (73, 76). Although a usefultechnique, harmonic imaging has limitations. Most importantly, owing to inherent

a

A

0.5 cm0.5 cm b

Figure 6(a) B-mode ultrasound image of the left kidney of a 9-week-old mouse (bar = 0.5 cm). (b) Thesame kidney as in (a) is shown following intravenous administration of ultrasound contrast andapplication of power Doppler. Notice that the kidney vasculature is visualized to the level ofthe interlobular vessels (arrows). A = aorta.

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Fundamental signal

Harmonic signal

Returning signal

Figure 7In harmonic imaging, anultrasound pulse of aspecific frequency(fundamental signal) isemitted by the transducer.The transducer receivesthe emitted frequency aswell as the harmonic. Thereturned signal is recordedas a combination of thefundamental andharmonic frequencies.

properties of the technique, there is typically a compromise that must be made be-tween image contrast and spatial resolution (77). Additionally, owing to nonlinearsound propagation, tissue can also produce nonlinear echoes, thereby reducing con-trast resolution (76).

To overcome these limitations, phase inversion imaging was developed (Figure 8)(77). Phase inversion involves multiple ultrasound pulses that are emitted sequentially,where the second pulse is inverted compared to the first. Linear reflectors will send

Transmitted Received Summation Result

Tissue

Microbubbles

Figure 8In a phase inversion scheme, multiple pulses are transmitted with differing phase. If the pulsesare reflected by a linear reflector such as tissue, echoes appear similar to the transmitted pulses(top). The ultrasound system sums the returned echoes, which negate one another. If thereflector of the phase-inverted pulses is nonlinear, such as with a microbubble (bottom), theinverted pulses will be altered upon their return. Therefore, their sum will not negate, and asignal is detected.

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the transmitted pulses back unchanged such that the sum of the paired waves willcancel and give no signal. However, nonlinear pulse pair responses will result in echochanges such that they no longer cancel when summed. Contrast-enhanced phaseinversion imaging has been used to subjectively assess neovascularity (78). In anotherstudy, contrast-enhanced phase inversion harmonic imaging of xenografted tumorswas shown to correlate with a semiquantitative scale of immunohistochemical stainingwith COX-2, a predictor of tumor angiogenesis (79).

Other multipulse contrast imaging schemes have been introduced, includingcontrast-specific sequence (e.g., CPS® from Siemens), which uses a combinationof changes in pulse phase and amplitude to selectively minimize tissue echoes andenhance ultrasound contrast agent echoes. With selective scaling of each receivedecho and addition of all echoes, tissue signals are rejected and nonlinear microbubbleechoes are retained (80, 81). An image registration feature is available on commercialultrasound systems, minimizing motion and improving the accuracy of region-of-interest (ROI) measures.

As noted above, ultrasound contrast agents can be destroyed when subjected toultrasound of appropriate frequency and sufficient pressure (30, 82, 83). A destructivepulse of ultrasound can be used to fragment the contrast agent within the region beingimaged (Figure 9). Lower pressure, higher frequency, nondestructive pulses can beused to visualize the replenishment of intravascular contrast agent into the imagingregion. The echo amplitude can be measured as the contrast agent refills the region,and the rate of return can be estimated. A destruction-replenishment method wasdescribed in 1998 by Wei et al. for the assessment of blood flow velocity in themyocardium (84). Continued research has assessed microbubble replenishment afteracute myocardial infarction (84–87) and in abdominal organs such as kidney and liver(88, 89). An adaptation to the Wei model of replenishment kinetics has recently beendescribed for single bolus intermittent contrast-enhanced power Doppler imaging oftumor perfusion in small animals (90). Nonlinear imaging techniques, such as CPS,can be combined with destruction-replenishment sequencing for the assessment ofregional blood flow (Figure 10).

IMAGING THE RESPONSE TO THERAPY

Combining a targeted imaging modality with lesion-directed therapy results in theability to determine several biologically relevant facts regarding the likelihood of apositive treatment response. Of particular interest are questions regarding whetherthe target is present, if the agent is reaching the target, and whether the intendedtarget is truly what is being treated. There are a variety of interesting biologicalprocesses amenable to the application of targeted ultrasound imaging for monitoringthe efficacy of drug delivery.

Our group has described a contrast-enhanced ultrasound technique combineddestruction-replenishment ultrasound with subharmonic phase inversion imag-ing to improve spatial resolution and to differentiate contrast echoes from tis-sue echoes (91). During the nondestructive imaging pulses, sound was transmittedfrom the transducer at a specified frequency, whereas the receive function detected

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Figure 9Ultrasound contrast agentis freely circulating in thevasculature. In adestruction-replenishmentscheme, a strongultrasound pulse destroysthe agent in the imagingplane. Low-pressurepulses are then used toobserve the contrast agentreturning to the imagingplane.

subharmonic frequencies of the original frequency. Subharmonic oscillations areuniquely generated by ultrasound contrast agents and not surrounding tissue (92),resulting in substantial subharmonic echoes from contrast agents within vessels andlittle to no signal from surrounding tissue. Quantitative parametric maps of flow ve-locity and overall integrated intensity were generated (93) and measures of perfusioncompared favorably with gold-standard techniques (94). This technique was used tomonitor the response of experimental tumors treated with antiangiogenic agents andidentified different levels of response to therapy (95).

CHANGES IN CELL MEMBRANE AND VASCULARPERMEABILITY

Electron microscopy has demonstrated that small holes produced within a cell mem-brane are associated with the collapse of a microbubble and production of a jet (96).

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a b

c d

Figure 10(a) A destructive pulsedestroys all contrast agentsin this rat kidney.(b) Contrast cadence pulsesequencing (CPS) visualizescontrast replenishment intothe kidney at 2 s whilenegating most signal fromregional tissue. (c) Contrastis refilling the kidneyvasculature at 12 s and(d ) 30 s after the destructivepulse. Time-intensity curvescan be generated from anypart of the kidney.

Depending on the ultrasound parameters, pores produced within a cell membranemay be short-lived, resulting either in cell death or successful introduction of an ex-ogenous material into the cytoplasm (5, 97). In addition to changing cell membranepermeability, the application of ultrasound to small vessels containing microbubblescan change blood vessel wall permeability, resulting in the extravasation of particlesinto the interstitial space (3, 4). This change in capillary permeability is dependent onbubble size, shell composition, and the ratio of capillary diameter to bubble diameter.Varying ultrasound parameters, such as the acoustic pressure and pulse interval, andphysical parameters, such as injection site and microvascular pressure, can maximizelocal drug delivery of microspheres. A pressure of 0.75 MPa at an ultrasound centerfrequency of 1 MHz is sufficient to produce capillary rupture in the exteriorized ratmuscle microcirculation (98). The ultrasound pulse interval affects both the numberof points at which extravasation is observed and the volume of material delivered,with both maximized at a pulse interval of 5 s (98). It is believed that maximizing thevolume of material delivered requires the replenishment of microbubbles into theregion between pulses. It also was shown that as capillary blood pressure increases,microsphere transport across the capillary wall increases.

APPLICATIONS: POTENTIAL IMAGING ANDTHERAPEUTIC TARGETS

Ultrasound contrast agents are uniquely suited for local drug delivery because theycan be forced to oscillate within a region with instantaneous dimensions as smallas hundreds of microns or as large as centimeters. One approach is to systemicallyinject the pharmaceutical simultaneously with an ultrasound contrast agent and tothen apply ultrasound to the target region. In smaller blood vessels, the oscillations of

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Figure 11(a) Ultrasound contrast agents are freely circulating in small vessels along with drug particles(blue). Once a sufficiently strong ultrasound pulse is applied to the area, the contrast agentexpands rupturing the endothelial lining. Drug is then able to extravasate. (b) Drug-ladenultrasound contrast agents are freely circulating throughout the vasculature. A pulse ofultrasound is applied and ruptures the contrast agent, thereby liberating the drug payload.Because ultrasound is only applied in the region of interest, drug is preferentially deliveredlocally. (c) Drug-laden ultrasound contrast agents baring surface ligands targeted to specificendothelial receptors are freely circulating. The ligand preferentially binds the ultrasoundcontrast agent in the target region, increasing local agent accumulation. An ultrasound pulse isthen applied liberating the drug payload.

the bubble will alter the vessel walls, allowing for extravasation of the pharmaceuti-cal agent (Figure 11a) (3, 99). The feasibility of extravascular drug delivery has beenevaluated by co-injection of microbubbles with particles and dyes (4, 99). Co-injectionof microbubbles with gadolinium followed by MRI has shown that gadolinium ex-travasates (100, 101). Alternatively, drug may be incorporated into the microbubbleand increased local delivery achieved by rupturing the microbubbles selectively inthe feeder vessels of the lesion (Figure 11b) (55, 102). However, these methods donot eliminate wash-out and systemic distribution of released drug by flowing blood.Successful demonstrations of reduced neointimal formation, endothelial transfection,and clot dissolution by microbubbles have been reported (103–107). Although thevolume of the microbubble payload delivered thus far has been small, delivery ofdrugs or genes across the blood brain barrier (BBB) is a promising application ofmicrobubble-based delivery, since few alternative methods can change BBB perme-ability to such a wide range of cargos (100, 108, 109).

As defined earlier, radiation force has been described as a method to push the mi-crobubbles toward the vascular wall prior to destroying them. The close proximity of

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the microbubble to the vascular wall will effectively paint the drug onto the endothe-lium upon rupture of the delivery vehicle. This method has resulted in depositionof a tenfold increase in fluorescently labeled oil to cells in vitro in comparison toultrasound alone (53).

More specific drug delivery may be achieved by attaching a ligand that is di-rected at a designated surface marker to the outside of the drug-laden microbubble.For example, endothelial surface markers are particularly attractive targets as certainmarkers are overexpressed in areas of angiogenesis and targeted microbubbles havebeen shown to adhere to these markers (110). Ultrasound can be applied locally totarget bound microbubbles resulting in delivery of drug selectively in the areas wherethe surface marker is expressed (Figure 11c) (111).

Thrombolytic Therapy

The first successful targeted ultrasound contrast agents were developed in the late1990s using avidin-biotin adhesion (112, 113). For in vivo imaging, a three-step pro-cess was developed (113). First, a biotinylated monoclonal antibody was administeredthat bound to fibrin within the clot. Avidin was administered, which bound the biotinon the monoclonal antibody. Finally, biotinylated ultrasound contrast agents weregiven, which bound the exposed end of the avidin molecule. This method of ultra-sound contrast targeting resulted in a fourfold increase in acoustic signal from clots(113).

More recently, Unger et al. have developed an ultrasound contrast agent,MRX408, targeted to activated platelets. This agent uses an alternative bindingmethod with an arginine glycine aspartic acid (RGD) molecule attached directly tothe surface of the contrast agent (114). RGD binds to glycoprotein IIB/IIIA receptorspresent on activated platelets. MRX408 has proven to increase the visibility of thrombiand better characterize the extent of the thrombus both in vitro and in vivo (115, 116).

Ultrasound has been shown to enhance thrombolysis with and without the addi-tion of microbubbles and typically in conjunction with intravenous administration ofthrombolytic agents (117–119). Effective thrombolysis with minimized treatment-associated hemorrhage has been demonstrated with an ultrasound frequency of1-2 MHz. Targeted or free microbubbles can be administered intravenously or di-rectly into the clot. The mechanism behind ultrasound-guided thrombolytic therapyinvolves the mechanical properties of the microbubbles themselves. When insonifiedat low frequency and high power, contrast agents expand and contract and can po-tentially break apart the clot. In addition, thrombolytic agents such as t-PA can beincorporated into the bubble and deposited into the thrombus when the bubbles areruptured. The combination of cavitation and thrombolytic agent delivery has provenmore successful for dissolving clots than ultrasound with contrast agents alone.

Anti-Cancer Therapy

For years, lipid-soluble anticancer agents have been incorporated into delivery ve-hicles so as to avoid systemic toxicity. As described above, it is now possible to

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incorporate hydrophobic agents into the lipid outer layer of the imaging microbubbles(53) or to attach hydrophilic molecules to the bubble shell (120, 121). Alternatively,it is possible to immerse a hydrophobic drug into an oil layer of acoustically activeliposomes (AALs) (111). Toxicity studies indicate a tenfold reduction in toxicity whenAAL-encapsulated paclitaxel was systemically administered to mice and compared tounencapsulated paclitaxel.

Angiogenesis is an essential process for the growth and spread of cancer (122).Endothelial cells and receptors associated with angiogenesis are readily accessible tointravascular agents such as ultrasound microbubbles. The placement of ligands ontothe surface of ultrasound contrast agents has already proven successful for imagingpurposes. Integrins, and particularly αvβ3, play an important part in angiogenesis withroles in cell adhesion, cell migration, and signal transduction (123). Lindner’s group(124) has conjugated monoclonal antibodies and RGD peptides with a high affinityfor αv-integrins to the surface of microbubbles using an avidin-biotin system. Inmouse models, ultrasound detected a greater signal from these bubbles in angiogenicareas where αv-integrins were upregulated.

Endoglin (CD105), the receptor for transforming growth factor, is a proliferation-associated, hypoxia-inducible protein that is highly expressed on angiogenic endothe-lial cells (125). Immunoscintigraphy using 99mTc-labeled monoclonal antibodies tar-geting endoglin has shown substantial uptake in tumors. More recently, a novelmethod for conjugating monoclonal antibodies specific for endoglin to microbub-bles has been described (35). Avidin was incorporated into the shell of microbubblesby sonication and then bound to the monoclonal antibody by biotin. Ligand directedaccumulations of microbubbles targeting endoglin were demonstrated in vitro (35).Given the ability to attach peptides and monoclonal antibodies to microbubbles, onecan envision targeted ultrasound agents for the imaging of tyrosine kinase receptorsfor vascular endothelial growth factor (VEGF) (126), fibroblast growth factor (FGF),and tissue inhibitors of metalloproteases (TIMPS) (127).

Atherosclerotic Therapy

One of the earliest indicators of atherosclerosis is the activation and attachment ofmonocytes to endothelial cells (128). This is mediated by the upregulation of leuko-cyte adhesion molecules (LAMs) such as intercellular adhesion molecule-1 (ICAM-1).In 1997, ultrasound contrast agents with albumin shells used for routine myocardialcontrast echocardiography were seen to have a slow transit time through the my-ocardium under certain pathologic conditions (129). These microbubbles preferen-tially adhered to endothelial cells expressing LAMs in vitro. Subsequently, ultrasoundcontrast agents bearing monoclonal antibodies targeted to ICAM-1 were shown tohave good binding efficiency in vitro and in vivo (130). Active targeting of inflam-mation using microbubbles was described by Villanueva et al. (131, 132) and others(27, 133–135), where endothelial cells activated during the inflammatory responsewere targeted using a microbubble. Takalkar et al. (135) used a parallel-plate flowchamber to determine the adhesiveness of microbubbles targeted with anti-ICAM-1to endothelial cells artificially activated with interleukin-1β. A 40-fold increase in

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microbubble adhesion occurred with targeted microbubbles compared with a non-targeted control. Microbubbles adhered at shear rates up to 100 s−1, which is char-acteristic of larger venules.

Other leukocyte adhesion molecules are up-regulated in inflammatory andischemia-reperfusion injuries. Of particular interest is P-selectin, which has beentargeted with ultrasound contrast agents in a mouse model of inflammation (133).Rychak et al. recently demonstrated targeted adhesion of deformable microbubblesto P-selectin (27).

Gene Therapy

One potential therapeutic application of targeted ultrasound contrast agents is forthe delivery of gene-based therapies. Nonspecific regional delivery of adenoviral andplasmid reporter genes has been accomplished using ultrasound-directed methods.More specifically, adenoviral or plasmid vectors have been incorporated into albuminbased ultrasound contrast agents and delivered to the myocardium using ultrasoundto destroy the microbubbles in the target region (136, 137). Microbubbles carryingplasmids encoding VEGF have been used to induce angiogenesis in the myocardiumof rats following ultrasound application (138). However, traditional microspheresare negatively charged and have a low efficiency for cell transfection of negativelycharged RNA and DNA molecules. Tiukinhoy et al. developed a positively chargedliposome with ultrasonographically detectable echogenic properties (139). Using anintravascular ultrasound system, they were able to deliver and detect luciferase geneexpression in HUVEC cells following ICAM-1-targeted, ultrasound-directed genetransfection.

Incubation of DNA and microbubbles can result in fusion of the DNA to the shell(140), thus facilitating co-injection. Early studies demonstrated gene delivery to themyocardium following intravenous injection of albumin microbubbles, incorporatingplasmid DNA on the shell, together with ultrasound (137). Subsequent studies havedeveloped techniques to incorporate DNA within lipid microbubble shells with sim-ilar local transfection after intravenous injection and insonation (56, 120, 136, 138,141–144). Although successful transfection has been reported using venous injection,a study comparing intravenous with intra-arterial injection of plasmid-containingmicrobubbles concluded that intra-arterial infusion was 200-fold more efficient inachieving local tissue transfection (54).

Delivery of therapeutic levels of drug or therapeutic gene delivery has yet tobe demonstrated with intravenous injection of clinically-relevant concentrations ofmicrobubbles. Cardiac gene transfection in the rat utilized 1 ml of intravenouslyinjected ultrasound contrast agents with a concentration on the order of 1 × 109

microbubbles/ml (136). Effective delivery of a therapeutic gene to the rat pancreaswas achieved with injection of 1 ml of microbubbles containing the gene within theshell, injected at a concentration of 5 × 109 microbubbles/ml (55). These studies em-ployed a dose much larger than that recommended for human imaging. The develop-ment of microbubble agents capable of producing successful transfection with a smalldose of intravenously injected microbubbles will be important for future translational

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studies. It is not clear, however, whether high concentration is required due to thelow payload capacity of the microbubbles or the need for a high concentration ofgas bubbles to be present when ultrasound is applied. Alternatively, interventionaltechniques in which a high concentration of microbubbles is injected within a muscleor artery to achieve local drug or gene delivery may be considered. Intramuscular in-jection of microbubbles and plasmids produced consistent local transfection in smallpreclinical studies (145). With intramuscular injection of plasmid and microbubbles,the injection of microbubbles alone enhanced transfection without ultrasound (146).Coinjection of plasmid DNA and microbubbles into the renal artery, when combinedwith transient vascular compression and ultrasound, has been shown to produce localgene expression in the kidney (147). Coinjection of plasmid DNA and microbub-bles into the cerebrospinal fluid, coupled with ultrasound, produced DNA transferinto the rat central nervous system (148). Tsunoda et al. (149) showed an order ofmagnitude increase in reporter gene transfection to the heart after local injection ofmicrobubbles and plasmid DNA to the left ventricle compared to injection via thetail vein.

CHALLENGES AND LIMITATIONS

Several important challenges face the field of targeted imaging and drug delivery. Themost difficult yet essential questions to answer include the following. Is the targetpresent in this individual lesion? Is the ligand specific for the target, or could it bindsome other site? Is the target abundant enough to bind detectable amounts of theagent? Will the drug payload remain locally or wash out into the systemic circulation?

Additional issues are more specific to ultrasound contrast agents as drug deliveryvehicles. First, ultrasound contrast microbubbles are relatively large compared withtraditional pharmaceuticals. Microbubbles are typically 1–10 μm in diameter. Tumorvessels are particularly permeable and often have large endothelial gaps (150); how-ever, contrast microbubbles are typically too large to exit the vasculature. This poses aparticular problem when trying to target receptors that may be present in the tumortissue rather than on the vascular endothelium. In a recent article by Wheatley etal., a nanoparticle ultrasound contrast agent (450 nm diameter) was described withgood acoustic properties (151). This agent resulted in good renal opacification inexperimental rabbits.

Although the circulation time of ultrasound contrast agents has increased in thepast several years, this is also a concern for ultrasound drug delivery. For example,the elimination half-life of Sonovue is 6 min (152). Uptake of Albunex occurs in theliver, lung, and spleen of rats and pigs, with 70% cleared from the bloodstream in3 min (153). If the agent is taken from circulation by the reticuloendothelial system,the circulation time may not be long enough to allow higher amounts of drug tobe delivered to the target region. Contrast agents are typically administered intoa peripheral vein so only a small amount of agent will pass through a tumor in agiven circulatory cycle. Multiple circulations are necessary to allow destruction ofenough agent to increase local concentration significantly. Polymer-shelled agentsmay provide a greatly increased circulation time (32).

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Although ultrasound microbubbles are relatively large agents, the amount of drugthat can be attached to the bubble surface or incorporated into the internal lipidlayer is a concern. In a study by Tartis et al., approximately 4 mg/ml paclitaxel wasincorporated into the internal oil layer of AALs (111). Currently, the recommendedsystemic paclitaxel dose approaches 175 mg/m2 (Taxol, FDA product insert, 1998),although it is administered at a low concentration to reduce toxicity. Since typicalstrategies for paclitaxel administration deliver less than 5% of the injected dose toa solid tumor, local microbubble delivery may allow regional delivery of a greatervolume than other strategies. Adding liposomes or other particles to the microbubbleshell also can substantially increase the efficiency of the vehicle (57).

Finally, the bioeffects associated with ultrasound contrast agents are not fullyunderstood and are strongly dependent on concentration and imaging parameters.These effects will need to be considered as clinical translation proceeds.

DISCLOSURE STATEMENT

The authors’ lab has a research collaboration with Siemens Medical Solutions (butdoes not receive funds) and joint grants with ImaRx Medical Inc.

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Annual Reviewof BiomedicalEngineering

Volume 9, 2007Contents

Cell Mechanics: Integrating Cell Responses to Mechanical StimuliPaul A. Janmey and Christopher A. McCulloch � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 1

Engineering Approaches to BiomanipulationJaydev P. Desai, Anand Pillarisetti, and Ari D. Brooks � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 35

Forensic Injury BiomechanicsWilson C. Hayes, Mark S. Erickson, and Erik D. Power � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 55

Genetic Engineering for Skeletal Regenerative MedicineCharles A. Gersbach, Jennifer E. Phillips, and Andrés J. García � � � � � � � � � � � � � � � � � � � � � � 87

The Structure and Function of the Endothelial Glycocalyx LayerSheldon Weinbaum, John M. Tarbell, and Edward R. Damiano � � � � � � � � � � � � � � � � � � � � � �121

Fluid-Structure Interaction Analyses of Stented Abdominal AorticAneurysmsC. Kleinstreuer, Z. Li, and M.A. Farber � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �169

Analysis of Time-Series Gene Expression Data: Methods, Challenges,and OpportunitiesI.P. Androulakis, E. Yang, and R.R. Almon � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �205

Interstitial Flow and Its Effects in Soft TissuesMelody A. Swartz and Mark E. Fleury � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �229

Nanotechnology Applications in CancerShuming Nie, Yun Xing, Gloria J. Kim, and Jonathan W. Simons � � � � � � � � � � � � � � � � � �257

SNP Genotyping: Technologies and Biomedical ApplicationsSobin Kim and Ashish Misra � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �289

Current State of Imaging Protein-Protein Interactions In Vivo withGenetically Encoded ReportersVictor Villalobos, Snehal Naik, and David Piwnica-Worms � � � � � � � � � � � � � � � � � � � � � � � � � � �321

Magnetic Resonance–Compatible Robotic and Mechatronics Systemsfor Image-Guided Interventions and Rehabilitation: A Review StudyNikolaos V. Tsekos, Azadeh Khanicheh, Eftychios Christoforou,and Constantinos Mavroidis � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �351

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SQUID-Detected Magnetic Resonance Imaging in Microtesla FieldsJohn Clarke, Michael Hatridge, and Michael Moßle � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �389

Ultrasound Microbubble Contrast Agents: Fundamentals andApplication to Gene and Drug DeliveryKatherine Ferrara, Rachel Pollard, and Mark Borden � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �415

Acoustic Detection of Coronary Artery DiseaseJohn Semmlow and Ketaki Rahalkar � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �449

Computational Anthropomorphic Models of the Human Anatomy:The Path to Realistic Monte Carlo Modeling in Radiological SciencesHabib Zaidi and Xie George Xu � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �471

Breast CTStephen J. Glick � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �501

Noninvasive Human Brain StimulationTimothy Wagner, Antoni Valero-Cabre, and Alvaro Pascual-Leone � � � � � � � � � � � � � � � � � � �527

Design of Health Care Technologies for the Developing WorldRobert A. Malkin � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �567

Indexes

Cumulative Index of Contributing Authors, Volumes 1–9 � � � � � � � � � � � � � � � � � � � � � � � � � � �589

Cumulative Index of Chapter Titles, Volumes 1–9 � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �593

Errata

An online log of corrections to Annual Review of Biomedical Engineering chapters(if any, 1977 to the present) may be found at http://bioeng.annualreviews.org/

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