Böhmer et al 2009

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European Journal of Radiology 70 (2009) 242253

Contents lists available at ScienceDirect

European Journal of Radiology

j o u r n a l h o m e p a g e : w w w . e l s e v i e r . c o m / l o c a t e / e j r a d

Ultrasound triggered image-guided drug delivery

Marcel R. Bhmer a,c,, Alexander L. Klibanov b, Klaus Tiemann c, Christopher S. Hall d,Holger Gruell a, Oliver C. Steinbach a

a Philips Research Europe, Biomolecular Engineering, HTC11, 5656 AE Eindhoven, The Netherlandsb Cardiovascular Division, Department of Medicine, Cobb Hall, University of Virginia School of Medicine, Hospital Drive, Cobb Hall RM 1026, Charlottesville, VA 22908-158, USAc Department of Cardiology and Angiology, University Hospital Mnster, Albert Schweitzerstrasse 33, 48149 Mnster, Germanyd Philips Research North America, Ultrasound Imaging and Therapy, 345 Scarborough Road, Briarcliff Manor, NY 10510, USA

a r t i c l e i n f o

Article history:

Received 13 January 2009

Accepted 14 January 2009

Keywords:

Ultrasound

Image guidance

MRI

Microbubbles

Temperature sensitive liposome

a b s t r a c t

The integration of therapeutic interventions with diagnostic imaging has been recognized as one of

the next technological developments that will have a major impact on medical treatments. Important

advances in this field are based on a combination of progress in guiding and monitoring ultrasound

energy, novel drug classes becoming available, the development of smart delivery vehicles, and more in

depth understanding of the mechanisms of the cellular and molecular basis of diseases. Recent research

demonstrates that both pressure sensitive and temperature sensitive delivery systems hold promise for

local treatment. The use of ultrasound for the delivery of drugs has been demonstrated in particular

the field of cardiology and oncology for a variety of therapeutics ranging from small drug molecules to

biologics and nucleic acids.

2008 Elsevier Ireland Ltd. All rights reserved.

1. Introduction

Theroleof medical imaging technologies in medical care is shift-

ing from a tool for diagnosis of a disease to being an integral part

of therapeutic interventions such as in image-guided treatments.

Stereotactic systems use images obtained before surgery, e.g., MR

and CT, foraccurateguidanceof a surgical toolto the target anatomy.

Instead of tissue removal, one can use high intensity focused ultra-

sound (HIFU) as a surgery tool. Using HIFU, energy can be focused

precisely to a small volume of interest. HIFU allows ablation of

tissue by local administration of thermal dosages. Image-guided

therapy offers the potential to direct therapeutic action precisely to

the point in the tissue where it is needed and not to other tissues.

When this is possible, a high and local thermal dose can be admin-

istered. Image-guided delivery using HIFU requires the integration

of imaging for diagnosis andtreatment planning anda therapy that

can be accurately directed and controlled by simultaneous imageguidance, resulting in less side effects.

The use of ultrasound for local hyperthermia was recognized

early as reviewed by Moyer and Clement [1,2]. Direct exposure

to therapeutic ultrasound produces irreversible cell death through

coagulative necrosis, and is currently being clinically evaluated in

breast, kidney, and liver tumors [3]. There is an increasing level of

literatureevidence [39] that demonstrateshow ultrasound energy

Corresponding author.

E-mail address: [email protected](M.R. Bhmer).

can also be used non-destructively for increasing the efficacy for

delivery of drugs and genetic material. Especially for chemothera-

peutic regimens to be successful in cancer treatment, the particular

drug must be effective in the tumor environment and administered

in doses that cause tumor eradication while keeping severe side

effects within acceptable limits, commonly called the therapeutic

window.

Performing minimally invasive therapy, such as ultrasound

mediated drug delivery (USDD), under image guidance requires

adequate definition of the region of interest and accurate compen-

sation for motion. Especially in the heart the feedback provided is

necessary to target the therapy accurately.The region of interestcan

be identified by detection of an abnormal morphology. Molecular

imaging holds promise to apply minimally invasive therapy in an

early stage of a disease as malignancies can be detected in an early

stage. Molecular imaging uses targeted contrast agents, which are

agents decoratedwith, for instance,antibodies or fragmentsthereofthat specifically interact with specific markers such as endothelial

markers of inflammation or angiogenesis.

New methods in ultrasound and magnetic resonance (MR)

provide higher resolution information in two and three spatial

dimensions, with acquisition and display occurring nearly in real

time. Computer image processing methods offer ways of clarifying,

highlighting, or detecting specific regions in tissue. Developments

in MR thermometryprovidea technical solution tofollow thedeliv-

ery of a thermal dose to a lesion. Fortreatment, a volumeof interest

inside a patient is delineated based on MR imaging, and subse-

quently heated by focused ultrasound. The tissue temperature is

0720-048X/$ see front matter 2008 Elsevier Ireland Ltd. All rights reserved.

doi:10.1016/j.ejrad.2009.01.051
http://www.sciencedirect.com/science/journal/0720048Xhttp://www.elsevier.com/locate/ejradmailto:[email protected]://dx.doi.org/10.1016/j.ejrad.2009.01.051http://dx.doi.org/10.1016/j.ejrad.2009.01.051mailto:[email protected]://www.elsevier.com/locate/ejradhttp://www.sciencedirect.com/science/journal/0720048X

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M.R. Bhmer et al. / European Journal of Radiology 70 (2009) 242253 243

mapped by MR thermometry and fed back into the control of the

ultrasound transducer to provide full temporal and spatial control

of the heating [1012]. Therefore, the technique gives more than a

feedback on the location of the region of interest, it also provides

information on a physical parameter, which can be used to guide

and control the therapy [13,14]. MRI can also be used to monitor

changes in the permeability of the vasculature, as shown by Treat

et al. [15]

Contrast ultrasound imaging, using microbubbles, also pro-

vides useful information for image-guided drug delivery. For these

agents, optimized detection algorithms are available on ultrasound

diagnostic imaging systems. With respect to therapy transducer

design, developments in electronic steering of the beam improve

the size of focal region and reduce grating lobes while maintaining

a small number of elements and a compact size.

In this review we focus on the use of ultrasound for therapy

and provide examples in cardiology and oncology. We will review

drug delivery vehicles based on temperature and pressure sensitive

systems. Such systems are either modified slow release systems

with a temperature sensitive component or contrast agents that

have been modified to include or enhance drug delivery. Develop-

ments in ultrasound and MRI imaging, and new agents to follow

and quantify drug release, will be described.

2. Ultrasound mediated drug delivery systemsequipment

design considerations

The equipment for ultrasound mediated delivery varies widely

dependent on the application and often on the clinical availabil-

ity of ultrasound imaging or therapy devices. The following section

will describe the components parts of ultrasound therapy devices

for drug delivery with a discussion on the relevant importance and

design limitations. We will then follow with a discussion of spe-

cific application requirements that depend on the target volume

within the subject. As diagnostic imaging systems are not designed

for therapy, which is in particular reflected in the focusing of the

ultrasound beam, we will not consider these studies here.

2.1. Signal excitation

The two primary physical mechanisms for activating particle

based therapies can be separated into heat- and pressure-activated

particles. In the case of heat-activated particles, typically the

goal is to deposit enough acoustic energy into the targeted vol-

ume to raise the temperature in order to release an encapsulated

drug. In these cases, the duty cycle of the acoustic signal is of

paramount importance and so the electrical excitation and ampli-

fication will consist of pure tone (long pulse lengths) signal sources

and high power amplifiers (class D op-amps are popular). In the

case of pressure-mediated release, the acoustic signal is often of

shorter duration and lower overall energy deposition than thatused in heat-release applications. Pressure based release uses par-

ticles, mostly microbubbles, which deposit or release a drug when

encountering a peak negative pressure beyond a particular thresh-

old (usually) between 0.5 and 5 MPa peak negative pressure. These

signals are quite short in duration but high in acoustic pressure. A

typical approach is to transmita short (

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244 M.R. Bhmer et al. / European Journal of Radiology 70 (2009) 242253

of the organs, and amount of peripheral blood flow for heat con-

vection among other factors are not simple linear scale factors. For

example, thedepth of thetreatment zone will determinethe choice

of ultrasonic frequency which canlead toinadvertentheating in the

overlying tissue in the case of heat-activated particles, which may

also cause release in areas not intended. Also, the choice of a lower

frequency to penetrate to deeper tissue can negatively impact the

heat deposition as the lower frequencies are not absorbed at the

same rate as higher frequencies.

2.4. Steering

In many applications for ultrasound mediated drug delivery, it is

advantageous to be able to treat a large volume of tissue with great

spatial resolution. Many approaches have been employed with two

major divisions: spatial movement of the transducer and electronic

steering of the focus of the transducer. The first approach is simply

performed by placing a single element transducer (or low element

count) on a translatable stage to allow large volumes of tissue to

be addressed. In many applications this is appropriate and cost-

effective, but it is not fast. Electronic steering is accomplished by

dividing a therapy transducer into multiple, individually address-

able parts. The electronic signalapplied to each element is retardedby a phase shift in such a way to control the location in the tissue

wherethe acoustic waves coherentlyinterfere. Such steering allows

for rapid (on order of milliseconds) changes in treatment location.

These arrays havebeen used indiagnostic imaging andareknown as

theclassof phasedarrays. In therapyapplications, theuse of phased

arrays has been more limited because of technical challenges as

mentioned in the following sections.

Several issuesmust be addressed when using arrays of elements

in therapeutic drug delivery. In particular, the size of the array can

often be large because of the needed pressures or acoustic energy.

As a result, when dividing into multiple elements, the element

count can be quite largeoften into the thousands in order to avoid

effects such as grating lobes. Grating lobes occur because of the

inadvertentphasecoherence occurringin theacousticfield in unin-

tended areas. This complication has implications for drug delivery

especially in cases where exposure or release of a drug in a sensitive

organ may lead to undesired side effects. Clever approaches have

been suggested to avoid these grating lobes without requiring a

large number of elements,includingthe use of sparse arrays,irregu-

larshapedand spaced elements,and limiting the number of phases

to be applied to the elements to simplify the driving electronics.

3. Pressure-mediated delivery

3.1. Ultrasound contrast agents

Microbubbles used as ultrasound contrast agents are tiny gas

bubbles, small enough to pass the lung capillary bed. To preventdissolution of the gas they have a shell made from a lipid, a pro-

tein or a biodegradable polymer. Lipid-shelled microbubbles are

used in clinical practice and have a monolayer of phospholipid.

An albumin-shelled agent, Optison, is also clinically available and

polymer-shelled agents have reached the end of phase III clini-

cal trials. Microbubbles are used for left ventricle opacification;

the ultrasound contrast between the blood in the left ventricle

and the myocardium is low and can be increased significantly by

intravenous injection of a small number of microbubbles. Typically

108109 microbubbles are injected for a diagnosticultrasound scan.

Microbubbles can be used to improve endocardial border delin-

eation and, thereby wallmotion abnormalities. In addition it allows

for analysis of myocardial perfusion, which further helps to iden-

tify the myocardiumat risk. Perfusion cannotonly be applied to the

Fig. 1. TEM picture of 2m (polylactide)-shelled microbubbles prepared by emul-

sification and freeze-drying technology.

myocardium but also to other tissues. As noted by Cosgrove [19] andSchneider [20], microbubbles can be used for dynamic detection of

macro and microvascular flow in many organs. Microbubbles are

also used to study the blood supply to the liver. Primary hepatocel-

lular carcinomas (HCC) are supplied by the hepatic artery. After a

bolus injection of ultrasound contrastagent, these lesions arehigh-

lighted by the perfusionof contrastagentbefore the rest of the liver

is fully perfused. Contrast liver imaging has been the subject of a

multi-center study and described by Lencioni et al. [21]

3.2. Thin-shelled and hard-shelled microbubbles

Microbubble agents can be classified as soft- or thin-shelled

and hard-shelled agents. Ultrasound contrast agents do not only

scatter ultrasound efficiently, they also react to low energy ultra-sound by emitting specific frequencies. Thin-shelled agents are

microbubbles having a lipid monolayer with a thickness of about

23nm. They undergo volume expansions and contractions that

generate an acoustic signal [22], of which non-linear components

give the most specific information for imaging [23,24]. As the shell

of these microbubbles is so thin, fluorinated gases are needed to

keep the microbubble stable for a sufficient time in the circula-

tion. Hard-shelled microbubbles have typical shell thicknesses in

the range of 20100 nm.An example of polymer-shelled microbub-

bles is given in Fig. 1. They hardly show volume expansions at low

acoustic pressure [25,26]. Nevertheless some of these agents do

generate acoustic signals as well without losing gas. Studies using

an extremely fast camera [27] have given first indications that they

often indent like a badly inflated ball, which is a way to conservetheir surface area and allow for a change in the volume [28,29]. At

higherpressures the microbubbles are destroyed showing dramatic

changes in the gas volume as shown in Fig. 2, where the activa-

tion of a polymer-shelled microbubble is given. Polymer-shelled

microbubbles do not need very hydrophobic gases to be stable in

circulation.

3.3. Targeted microbubbles

The use of microbubbles is currently being extended to targeted

imaging and drug delivery applications. For molecular imaging

applications, the shell is coated with specific ligands. A typical

example is the targetingof endothelial markers of angiogenesis and

inflammation [30,31].

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Fig. 2. Series of events upon insonation of polymer-shelled microcapsule, frames 1416, 20, 40, 60 of a movie recorded at 15 and 2.25 MHz; MI = 1. Ultrasound switched on

at frame 15. Movie taken on the Brandaris128 camera at Erasmus Medical Centre and Twente University {Chin, 2003 #156).

In the early proof-of-concept phaseof research,targeting ligands

are conveniently bound to the shell via biotinstreptavidinbiotin

bridges, see Fig. 3. Biotinylated lipids and biotinylated biodegrad-

able polymers [32] used for microbubble shell preparation are

available or can be synthesized. Microbubbles targeted to vascu-

lar endothelial growth factors and selectins have shown strong

enhanced ultrasound images in the areas of upregulation of these

markers in the vasculature [33].

Direct coupling of targeting ligands to the microbubble shell

using peptide bond formation chemistry [34], see Fig. 3, is nec-

essary at the next step towards clinical trials, when the presence

of non-endogenous proteins, such as avidin or streptavidin, is not

desired.

Initial model system studies [35] showed that biotinylatedmicrobubbles can be targeted to avidin-coated surface in vitro,

and ultrasound imaging of these targeted bubbles was success-

ful. Most of the experimental targeted ultrasound imaging efforts

have been focused on the various in vivo animal models, from

thrombus targeting [36] to the ultrasound imaging of a variety of

molecules upregulated on vascular endothelium. Anti-P-selectin-

Fig.3. Inclusionof drugs in lipid-shelled microbubbles. (A)Lipid-shelledmicrobub-

bleconsistingof gas encapsulated bya lipid monolayer,a fractionof thelipidcan be

pegylated (not shown). (B) Lipid-shelled microbubble with an additional oil-phase

to increase the reservoir size to incorporate hydrophilic drugs {Unger, 1998 #122}.

(C) Lipid-shelled decorated with liposomes via biotin streptavidin bridges.

antibody-carrying microbubbles have been successfully used for

targeted ultrasound contrast imaging in the areas of TNF-induced

inflammation or ischemia-reperfusion injury [37]. Detection of

ICAM-1 upregulation in transplant rejectionmodel was achieved by

targeting microbubbles with biotinylated anti-ICAM-1 antibodies

[38]. A large set of studies was conducted at imaging angio-

genic endothelium, via biotinylated antibodies against v3 in thetumor vasculature setting [39] as well as therapeutic angiogene-

sis [40]. Tumor vasculature status can be evaluated by targeting

streptavidin-carrying lipid microbubbles decorated with biotiny-

lated antibodies against VEGF receptor 2 [41]. The ease of use of

ultrasound imaging allows comparative targeted imaging of two

markers in the tumor vasculature of the same animals, for instance

endoglin versus VEGF receptor [33].At this time, covalent coupling methods, lacking avidinbiotin

scheme, aregaining wider acceptance, showing the successful cou-

pling of small ligands [42] or antibodies [43] with good yield. This

covalent approach will be more applicable in the clinic.

Binding of antibodiesto the vascular endothelium targets is very

strong and selective, but the formation of the bond between the

antibody and antigen is, typically, relatively slow. As the microbub-

bles at the target surface experience shear, especially in the fast

(arterial) flow, the relatively long time required to obtain firm

binding might not be sufficient for the antibody; for instance, in

the flow having wall shear stress over 2 dyn/cm2, anti-P-selectin

antibody-targeted microbubbles are not accumulating at the tar-

get efficiently [44]. To achieve leukocyte adhesion to the inflamed

endothelium, nature has a set of fast-binding ligands on the leuko-cyte membrane, such as PSGL-1 glycoprotein, that binds to P-

and E-selectin. Glycosulfopeptide-carrying microbubbles were tar-

geting P-selectin-coated surfaces in fast flow conditions quite

successfully [45] A simple variant, essentially a portion of the same

PSGL-1molecule, sialyl Lewis X, can be immobilized on microbub-

bles. Microbubble targeting via this ligand can be assisted by

co-immobilizing the antibody on the bubble, so the rapid attach-

mentof microbubblesto the target is aidedby firmantibodybinding

[46]. An alternative is to increase the ligand concentration on the

microbubble surface, e.g., by using polymeric version of sialyl Lewis

X, which is available commercially, polymeric sialyl Lewis X, is

capable of firmbut rapid cooperative multipointbindingwith the P-

selectin target surface, and providesefficientmicrobubble targeting

[47].

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Fig. 4. Polymer-shelled microbubbles (A) have a thicker shell into which drugs can be incorporated directly, hydrophilic drugs can be incorporated with a double emulsionmethod [60] (B). Half-oil filled polymer-shelled microbubbles (C) give an additional liquid reservoir into which hydrophobic drugs can be incorporated [28], the drug can

either be in solution or precipitated. Multilayer constructs (D) or the attachment of liposomes (E) is also possible with polymer-shelled microbubbles.

3.4. Therapeutic use of microbubbles: sonoporation

For ultrasound induced drug delivery based on microbubbles

two approaches are distinguished, see for instance the review by

Hernot and Klibanov [5] where the distinction is made between

co-administration of drugs and microbubbles and drug-loading of

the microbubbles themselves. In the case of co-administration the

function of the microbubbles is to enhance the permeability of

the endothelial wall. This can either be affected by rupture of the

endothelial wall, leading to extravasation of relatively large enti-

ties such as red blood cells and polymer particles [48] or in a more

subtle way, at lower pressures, by microbubbles causing temporaryopening of cell membranes.Many of the properties of cell mem-

branes are shear dependent. Marmottant and Hilgenfeldt [49,50]

demonstrated that the oscillations of microbubbles induce local

shear stress by altering the flow of liquid near the cell surface. Van

Wamelet al.[51,52] haveshown the deformation of cell membranes

in the presence of an oscillating microbubble directlyusing an ultra-

fast camera. Sonoporation is the term that is used to describe the

formation of pores by ultrasound. If the pores are too large they

cannot reseal leading to cell death, however, if the pores can seal

again they will stay open for a time and, in principle allow passage

of therapeutics, such as radionuclides [53] or plasmid DNA (see ref.

[8] for an overview). Mehier-Humbert et al. [54] investigated the

percentage of GFP positive Matt-B III cells following plasmid DNA

delivery using lipid- and polymer-shelled microbubbles.A great challenge in sonoporation is to open the blood brain

barrier in a reversible way. Treat et al. [15] and Hynynen et al. [55]

have shown that pores made in a rabbit brain close again in about

6 h. They also demonstrated that contrast enhanced MRI is a very

suitable method to follow sonoporation. The method has been used

to deliver doxorubicin across the blood brain barrier in rats, as well

as to deliver genes under MRI guidance.

As stressed in this paragraph the combination of microbubbles

and ultrasound have an effect on the properties of adjacent cells. A

number of authors have found that this effect extends to tumor

growth. For instance Miller and Song [56] reported that tumor

growth of renal carcinoma in mice in the presence of Optison and

ultrasound is reduced. Damaging tumors by ultrasound could have

an effect on metastasis because of the disintegration of the tumor.

Miller and Dou [57] investigated the enhancement of lung metas-

tasis from an implanted mouse melanoma tumor after application

of ultrasound in the presence of microbubbles. At high pressure

(5MPa) and a 1 Hzrate toavoidheating morelungmetastases were

indeed found in the presence of microbubble than in their absence.

However at lower pressure (2 MPa) no enhancement was found. In

the absence of microbubbles an elevated level of metastasis was

already found in at a peak negative pressure of 5 MPa.

3.5. Incorporation of drugs and genetic material in microbubbles

Instead of co-injecting drugs and microbubbles, microbubblescan also be modified to contain drugs [58,59] or DNA [6062]. The

advantage, as explicitly shown by Lentacker et al. [61] in an in

vitro setting is that the therapeutic molecule is close to where the

acoustic action is, and therefore opens the opportunity for a dose

reduction while maintaining its therapeutic efficacy.

If drugs have sufficient affinity for the lipid monolayer they can,

in principle, be incorporated directly in lipid-shelled microbub-

bles (Fig. 4A). However, as the monolayer is very thin, the amount

that can be incorporated is extremely low. Unger [59] has added

an additional oil, triacetin, to make a thicker hydrophobic layer

to increase the incorporation of paclitaxel, as shown schemati-

cally in Fig. 4B. This approach is limited to hydrophobic drugs; a

more general applicable route is to attach liposomes or lipoplexes

to the microbubble (Fig. 4C). Another option is to bind the drugto the outside, for instance in a form associated with lipids, such

a drug carrying liposomes or lipoplexes: positively charged lipid

complexed to negatively charged nucleic acid [61]. They have

demonstrated high transfection efficiency by showing luciferase

activity in vitro. Attachment of liposomes can allow the incorpo-

ration of both hydrophilic molecules into the aqueous core of the

liposome or hydrophobic drugs into the lipid bilayers. Finally mul-

tilayer technology [63] can be used to deposit large therapeutic

molecules layer by layer on a microbubble surface.

In principle layer-by-layer deposition or attachment of lipo-

somes is also possible for polymer-shelled agents, see Fig. 5.

However, different routes are also available to prepare drug loaded

polymer-shelled agents. Emulsification of a polymer solution con-

taining a carrier solvent and an alkane in an aqueous phase is a

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Fig. 5. Schematic representation of targeting by bioti-streptavidin bridges (left) and by NHS coupling (right).

starting point for making polymer capsules. The carrier solvent

is removed from the emulsion droplets and the polymer phase

separates from the shell closing in the alkane in the center. The

alkane is subsequently removed by freeze-drying. This prepara-

tion route is suitable for hydrophobic drugs. Fabrication techniques

starting from monodispersed emulsions, such as a technique based

on submerged inkjet-printing [64] allow for a precise incorpora-

tion of a known amount of drug per particle and therefore improve

the control over the doseage. Direct emulsification techniques can

also yield a narrow size distribution with a maximum between 1

and 2m and >95% of the microbubbles smaller than 3m [28].Hydrophilic drugs cannot be incorporated directly in an emulsi-

fication procedure. One possibility is to employ double emulsiontechnology. In a double emulsion, an aqueous phase, containing

the drug, is first emulsified in an organic phase (containing the

shell-forming polymer), and this first emulsion is subsequently

emulsified in a second aqueous phase. Poly-lactide-co glycolide

microbubblescontainingplasmidDNA have beenpreparedthis way

[60]. Gene delivery from these microbubbles was shown in rats and

also an effect on tumor growth was demonstrated.

An alternative way to prepare polymer-shelled microbubbles is

spray-drying [65] Hydrophobic drugs can in principle be included

in the spray. In the preparation method of Palmowski [66,67], the

starting point is not the polymer but a monomer. As the polymer-

ization reaction leads to the shell formation directly, incorporation

of a drug at this stage is difficult.

Recentdelivery experiments with a new, emerging class of genetherapeutics, small interfering RNA (siRNA), show promise to over-

come the inherent in vivo delivery obstacles of nucleic acids in

general and siRNA in particular, such as rapid excretion via the

liver, serum instability, non-specific distribution, tissue and cell

barricades [68]. The major limitation for the use of siRNA, both in

vitro and in vivo, is the inability of naked siRNA to passively diffuse

through cellular membranes dueto thestrong anionic chargeof the

phosphate backbone and consequent electrostatic repulsion from

the anionic cellmembranesurface.To deliversiRNA withmicrobub-

bles, siRNA was either directly attached to the microbubble surface

or simply mixed withmicrobubblesprior toadministration.In some

studies these vehicles showed enhanced transfection efficiency

both in vitro and in vivo. Further they provide a better protection

against degradation by serum nucleases. [60,69,70]

3.6. Targeted imaging and drug delivery with microbubbles

Microbubbles typically do not exhibit long circulation times.

The ReticuloEndothelial System takes them out of the circula-

tion and contrast is observed over time periods of about 20 min

in humans and a few minutes in mice and rats. The limited cir-

culation has consequences for targeted imaging and therapy as

compared to nanomedicine formulations. To achieve a long circu-

lationtime, small particles, around100200nm, have to be chosen,

and their surface charges have to be screened, for which espe-

cially poly-ethylene glycol is used. Although microbubbles can be

surface-modified with PEG, their size is optimized for acoustic

activity and therefore in the micron range. Fortunately the imag-ing of microbubbles is very sensitive and as the non-adhering

microbubbles disappear rapidly from the blood stream only the

few remaining adhering microbubbles can be imaged. Secondly

the acoustic signature of a microbubble differs, it shifts to lower

frequency, if the bubble adheres [71,72]. To exploit this for imag-

ing, however, monodisperse microbubbles are needed [73]. Finally

the use of radiation forces has been explored. These are long, low

amplitude acoustic pulses that drive the bubbles to the vessel wall

[7476]. Instead of long circulating small agents that pass by the

region of interest manytimes and may adhere once they pass close

to the vessel wall, microbubbles can be driven to the vessel wall

actively.

Compared to other imaging modalities targeted contrast ultra-

sound has the advantage that imaging can be performed relativelyfast at a high sensitivity. The microbubblesize, however, also brings

a drawback as it is more subject to shear forces in the blood flow

[32]. Although this has only be shown in a flow cell, the effects

of shear flow and microbubble displacement upon application of

ultrasound are aspects that need further study before ultrasound

imaging with targeted microbubbles can be used for more than

qualitative purposes.

The limited circulation time of microbubbles also has con-

sequences for their use as drug delivery vehicles. Treatment

preferably takes place shortly after injection of microbubbles and

will be restricted to well-perfused areas. Repeated injections and

ultrasound treatments are normally used to evaluate the effect of

ultrasound triggered release of drugs from microbubbles in terms

of tumor growth reduction.

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An approach to increase circulation times and aid extravasa-

tion is using nanoparticles, however, this will come at the expense

of decreased imaging possibilities. Nanoparticles, such as micelles

made of blockcopolymers, have been used forultrasound mediated

drug delivery and are reviewed by Husseini and Pitt [6]. In most

studies, the frequencies used are much lower than in those used in

combination with microbubbles. The mechanism of drug delivery

is also in this case related to cavitation. Rapoport [77] created dox-

orubicin containing polylactide nanoparticles, which, at least for a

fraction of them, contain perfluoropentane. Perfluoropentane has

a boiling point at 27 C, but in the form of emulsion droplets it is

superheated as shown by Giesecke and Hynynen [78]. Therefore

at body temperature they can still be in the liquid state and phase-

convertedby a trigger,such as an ultrasound pulse. The doxorubicin

containing particles are so small that they escape the vasculature

in a tumor by the enhanced permeation and retention effect and,

when exposed to ultrasound, cause pronounced tumor regression.

Long time imaging of remaining gas bubbles in tumor tissue was

possible.

4. Temperature sensitive drug delivery systems

While thermal ablation requires a substantial thermal dose toinduce tissue necrosis, a more subtle temperature increase can

be used to support treatment with conventional chemotherapeu-

tics and drug delivery systems [79]. Mild hyperthermia enhances,

for example, extravasation of drug loaded liposomes like Doxil

[80,81] or enhances anti-angiogenic treatment [82]. Hyperther-

mia can also increase local drug concentrations in conjunction

with temperature-induced drug delivery [83]. Temperature sensi-

tive drug delivery systems were already explored in combination

with hyperthermia induced by radiofrequency (RF) [84] magnetic

particles [85], or byheatingwithlightin theinfrared regime [86,87].

Only recently, some of these temperature sensitive drug delivery

carriers were explored in combination with ultrasound induced

drug delivery [88]. The more efficient uptake of drug delivery sys-

tems in tumors at elevated temperatures, together with the localtemperature triggered release of drugs, makes ultrasound induced

drug delivery a very promising field.

Temperature sensitive drug delivery systems can be designed

following two different approaches. One class of agents is based on

amphiphilic temperature sensitive polymers showing a lower crit-

ical solution temperature (LCST) in aqueous solution. The second

class of temperature sensitive drug delivery carriers is based on

liposomes. Here, lipids are used that show a phase transition above

body temperature. Upon passing the phase transition temperature,

the liposomal bilayer becomes leaky for drugs encapsulated in the

inner lumen of the liposome.

4.1. Polymer-based systems

Temperature sensitive polymeric drug delivery systems are

usually based on polymers that undergo upon heating a phase

transition associated with a change in polymersolvent interac-

tion [8991]. The solvent properties change from a good solvent

at temperatures below the LCST to a poor solvent at temperatures

above, leading to a morphology change from an extended random

coil toa collapsedchain(Fig. 6a). Most polymersstudied in this con-

text are based on N-isopropylacrylamide (NIPAAm) (Fig. 1b). The

LCST behavior of this polymer is due to a loss of hydrogen bonding

between the amino-group and surrounding water, and increased

hydrophobic interactions of the N-isopropylgroups above the tran-

sition temperature

Different strategies were followed to design temperature sensi-

tive drug delivery systems based on above mentioned temperature

Fig. 6. (a) Phase behavior of poly(N-isopropylacrylamide)-based polymers (b) in

water. At temperatures below the LCST the polymer is water soluble (random coil

configuration), while the chain collapses at temperatures above the LCST.

sensitive polymers. One strategy pursues micelles formed from

amphiphilic diblocks made up from a hydrophobic inert block-

polymer, like polylactic acid, polystyrene, etc. and a PNIPAAm

block [92,93]. The micelle can be loaded with drugs. Below the

LCST, this diblock self-assembles into micelles with a hydrophilic

PNIPAAmcorona. Heating induces a hydrophilichydrophobictran-

sition of the PNIPAAm block polymer, leading to a destabilizationand morphology change such as aggregation of the micelles. The

latter can significantly enhance drug release compared to tem-

peratures below the LCST. The LCST can be fine-tuned in a wider

temperature range by designing end-functionalized NIPAAm-based

polymers,copolymer or block polymers [94,95]. However, the com-

plex dependence of the LCST on intramolecular hydrogen bonding

and electrostatic interactions and interactions with water makes

the LCST susceptible to pH, ionic strength of the solvent and inter-

actions with other molecules. The advantage of delivery systems

where the drug release can be fine-tuned with respect to tempera-

tureand pH comeswith thedisadvantagethat drugreleasebecomes

also more complex and difficult to control in vivo as the LCST in

vitro and in vivo can significantly differ.Though many systems were

investigatedin vitro,little work hasbeen done in preclinicalstudiesin general or in combination withultrasound induced hyperthermia

in specific. One of the few preclinical studies exploits the interde-

pendency of pH and temperature to enhance drug delivery in the

more acidic environment of a tumor [96]. Other preclinical studies

showed the feasibility of temperature-induced drug delivery using

temperature sensitive polymeric micelles [97,98].

4.2. Temperature sensitive liposomes

Liposomal drug delivery systems are a well-studied field and

found their applications in cancer therapy in the clinic [99]. Lipo-

somes can be loaded with different hydrophilic drugs in the inner

water compartment. A particularly high drug payload is achievable

with drugs like doxorubicin or daunorubicin that precipitate in aninner lumen loading mechanismbasedon a pHgradient. Theconve-

nient method of drug loading and the achievable high drug payload

probably explain the dominant role of these drug delivery systems.

Drug release from conventional liposomal formulations is usually

diffusioncontrolled, showing little differencein drugreleasekinetic

at body temperature compared to temperatures slightly above in

the hyperthermia regime. The concept of drug delivery using tem-

perature sensitive liposomes was introduced by Yatvin et al. more

than 25 years ago using cis-platinum as a drug compound [100].

The field of liposome-based drug delivery under hyperthermia was

reviewed several times [79,101].

Temperature sensitive liposomes (TSL) are composed of lipids

thatshow a melting transition of the acyl-chains in the bilayer[102].

TSLs show a strongincrease indrug release around themeltingtran-

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Fig. 7. Temperature sensitive liposomes containing a drug (red dots) and an imaging or contrast agent that allows visualizing and quantifying the drug delivery process.

sition temperature Tm of the lipids associated with the formation

of transient pores in the bilayer [103,104]. The release temperature

andthe release kinetic can be controlledby choosing suitable lipids

with a Tm in the desired range of around 14 K above body temper-

ature and the incorporation of lysolipids [104,105], which induce

efficient pore formation after the melting transition of the mem-

brane. Temperature-induced delivery of doxorubicin encapsulated

in TSLs was thoroughly investigated in preclinical studies and is

now in clinical trials [106].

4.3. Temperature triggered release

In all temperature-induced drug delivery application, the chal-

lenge remains to quantify the amount of release drugs and

eventually control the release quantitatively. The achieved drug

concentration within a treated lesion can be evaluatedin preclinical

studies using standard analytical means, however, imaging tech-

niques such as nuclear imaging or MRI offer the advantage of being

applicable to humans in a clinical setting. Fig. 7 shows the under-

lying idea of incorporating a drug and contrast or imaging agent

inside a thermo-sensitive liposome. Upon heating the drug and

the contrast agents are released. The observable contrast originates

from a change of the local environment of the release contrast

agents, which correlates with thedrug release. Examples are T1 MRcontrast agents that are incorporated inside the liposome at high

concentration, and provide a strong signal once they are released

from the liposome.

Nuclear imaging is the method of choice to quantify the amount

of drug delivery systems accumulated in the lesion [107], or even-

tually also the released drug within the tissue. The disadvantage is

therequired radiolabeling of thedrug and/ordrug carrier andprob-

lematic integration of nuclear imaging in the treatment protocol as

a standard technique. Thus, nuclear imaging will probably keep its

role forvalidationin a research phase butwill notget a majorrole in

image-guided drug delivery. More promising is MR imaging of drug

release, especially in the combined setting of an HIFU/MRI system.

Here, MR contrast agents can be incorporated in the drug delivery

carriers that provide a signalchange upon drug release. The valueofthis approach was shown using for example temperature sensitive

liposomes filled with Gd-based T1 contrast agents or manganese

based agents [86,87,102,108110].

Temperature sensitive liposomes can also be designed by con-

jugating PNIPAAm polymers to the liposomal membrane [111,112].

Below the LTCS, the hydrophilic PNIPAAm polymer extends into

solution and stabilizes the liposome. Upon heating above the LCST,

the hydrophilic to hydrophobic transition leads to a destabiliza-

tion of the liposome and in vivo to a very different interaction with

cells.

No matter which approach is used, all temperature sensi-

tive drug delivery systems can be prepared in the size range of

10300nm, which presentsa majordifferencewith respect to pres-

sure sensitive micro bubbles. As most microbubbles stay within

the vascular system, released drugs still need to extravasate, possi-

bly aided by sonoporation to reach the target tissue. Blood flow

in the capillaries may carry away the released drug, diminish-

ing the effect high local drug concentration. Here, the smaller

size of temperature sensitive systems potentially offers the advan-

tage that the place of release can be chosen. At early times after

administration, drug release can take place in the microvascular

system when passing through a tissue with elevated temperature.

At later time points after administration, the drug delivery car-

rier may have accumulated first in a lesion either by passive oractive targeting mechanisms (e.g., EPR effect or using specific tar-

geting ligands). Drug release in the interstitial space is triggered

subsequently by heating the lesion using, for example, ultrasound.

However, alsoin the microvasculaturean increasedconcentrationis

found. For example, studies with doxorubicin loaded TSLs revealed

a 30-fold peak concentration of doxorubicin in the microvascula-

ture. Pronounced effects on tumor growth are also found if the

heat-treatment is given shortly after injection [108], therefore we

can only conclude that effective mechanism needs further study.

Treatment response may be a result of anti-vascular effects and

anti-neoplastic mode of action [101,113].

Future work will aim at designing new temperature sensitive

drug carriers that show a sharp and rapid drug release at temper-

atures slightly above body temperature. However, as much workwill be needed to optimize the treatment protocol with respect to

timing of injection, rate of extravasation, drug release kinetic, and,

finally, the timing of ultrasound induced hyperthermia itself with

existing drug delivery systems. The possibility of controlling the

drug delivery in real time under image guidance and quantifying

the drug release using, for example, MRI will be essential to estab-

lish a treatment protocol that is superior to todays approach with

better therapeutic value.

5. Therapeutic applications of ultrasound triggered drug

delivery

5.1. Cardiology

Ultrasound contrast agents have been approved for diagnostic

purposes in the filed of cardiology to better visualize the left ven-

tricle. Therefore many ultrasound contrast agent mediated drug

delivery studies have been performed in the heart as recently

reviewed by Mayer and Bekeradjian [8]. The paper summarizes

the field of gene delivery using ultrasound mediated delivery tech-

niques and gives examples of studies where the expression of

reporter genes was, in rats and mice, enhanced with a factor 3300

using ultrasound frequencies around 1 MHz. Also for therapeutic

genes an overview of a number of studies is given. The majority

of work in the cardiology field has focused on gene delivery but

the ultimate therapeutic aim is to induce angiogenesis and car-

diac repair. Side effects by low molecular weight drugs are less

of an issue in cardiology compared to oncology and have, there-

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fore, received less attention in the cardiology research on triggered

delivery.

As described previously, gas-filled microbubbles are an impor-

tant ultrasonic contrast agent used to enhance edge detection and

evaluate myocardial perfusion [114,115]. It was shown in preclinical

experiments that ultrasound mediated drug delivery can directly

enhance the expression of adenoviral vectors and plasmid DNA to

the heart [70,116,117,118].

USDD can successfully deliver plasmid DNA to myocardium. In

optimization experiments, the levels of reporter gene luciferase

expression were similar to that obtained using adenovirus but

without the profound liver uptake associated with adenovirus

[16].

In addition USDD has also been applied for organ-specific deliv-

ery of other bioactive agents [69] and could facilitate thedelivery of

protein therapeutics to ultrasound-accessible organs whilekeeping

systemic concentrations and side effects low. Vascular endothe-

lial growth factor (VEGF) bound to albumin microbubbles was

delivered to the heart using USDD. A more than 10-fold increase

of cardiac VEGF uptake was seen compared with systemic VEGF

administration [118]. Some preclinical studies have demonstrated

that ultrasound alone can facilitate uptake of various substances

including biologics.

Striking are the examples of Kondo et al. [18] to treat an acutemyocardial infarction resulting in enhanced angiogenesis and lim-

itation of the infarct size. Also, Korpanty et al. [119,120] have

shown increased density of arterioles and capillaries after ultra-

sound mediated gene transfer.

5.2. Oncology

In the last decades cancer has moved from a deadly to a man-

ageable chronic disease. Cancer in the breast, prostate, liver and

other organs can be imaged quite accurately with diagnostic ultra-

sound [121] and ifthesecarcinomas canbe segmented, targeted and

treated with therapeutic ultrasound, a new non-invasive, blood-

less approach to the treatment of such diseases can be developed.

However, after surgical removal and/or treatment with radiation orHIFUof a primary tumor,management of the residual tumor includ-

ing metastasis is typically carried out using a variety of systemic

therapies that include small organic molecules, and increasingly

innovative therapiessuch as biologics and the emerging siRNA ther-

apeutics [122,123].

For advanced tumor stages, chemotherapy remains the treat-

ment of choice. Despite the fact that such anticancer agents have a

very effective tumor killing potential in vitro and in animal cancer

models, they often fail in patients as they are unable to reach all

tumor cells that are able to regenerate the tumors [124] and, there-

fore,chemotherapy is rarelycurative butrather palliative,especially

for solid tumors [125].

The microenvironment of a tumor is critical in tumor initiation

andpromotion, andthere is increasing evidence that this may be animportant factor in developing therapeutic approaches [126]. The

tumor microenvironment, or stroma, influences the growth of the

tumor and its ability to progress and metastasize. It also can limit

the access of therapeutics to the tumor, alter drug metabolism and

contribute to the development of drug resistance. As opposed to

normal tissues, blood vessels in tumors are leaky [127] and vascu-

lature is less spatially organized [128], resulting in the abnormal

function of vessels [129]. The combination of impaired blood flow

through blockage by neoplasmatic tumor tissue, a leaky vascula-

ture, and lackof functional lymphatics leadsto increasedinterstitial

fluid pressures. In addition, the plasma to interstitial gradient of

osmotic pressure in tumors is also generally reduced limiting the

extravasation and creating a major obstacle against delivery of ther-

apeutic agents [130,131]

Furthermore, diminished oxygen delivery and hypoxic condi-

tions cause reduced efficacy of radiation therapy, high level of

metabolic products (e.g., carbonic and lactic acid), lower extracel-

lular pH and may potentially affect the cellular uptake of some

drugs.

Hyperthermia mediated liposomal drug delivery has shown

promise for enhancing local drug deployment while minimizing

drug distribution outside targeted tissues [101] and is currently

being applied clinically in the treatment of various types of cancer.

In cancer therapy, the studies, recently reviewed by Frenkel [3] on

temperature sensitive delivery vehicles, are more advanced [106],

however, radiofrequency ablation is the heating method, which is

more invasive, and has less well-defined spatial temperature con-

trol than ultrasound.

The clinical relevance of such controlled and triggered release

concepts for drug delivery systems having been demonstrated,

research in this area focuses currently on optimization of cell

specific targeting. These more advanced targeted nanocarriers in

general have clearly shown their potential in various animal tumor

models and await clinical application.

A more novel approach is to use gene therapy in cancer treat-

ment. A crucial requirement for gene therapy is tight control of

transgene expression, both spatial and temporal to enhance the

spatial targeting and efficiency of gene delivery. Tissuespecific pro-moters may also be used to limit transgene expression to targeted

tissues, and in that way add a layer of targeting and safety to gene

delivery procedures [132].

Therapeutic effects could be demonstrated in vivo with var-

ious targeted nucleic acid formulations, such as tumor-targeted

DNA plasmids expressing p53 or tumor necrosis factor alpha, small

interfering RNAs knocking down gene expression from tumor spe-

cific chromosomal translocations or gene expression of tumor

neoangiogenic processes, as well as double stranded RNA poly

inosinecytosine which triggers apoptosis in targeted tumor cells

[133].

In a wider sense, gene therapy is experiencing an unprece-

dented renaissance through the emerging field of the novel,

innovative drug format of small interfering (siRNA). Since thediscovery that double stranded (dsRNA) can specifically inhibit

expression of homologous genes, RNA interference (RNAi) has

become one of the most widely used methods for studying

loss-of-function phenotypes in model organisms and is increas-

ingly used across the whole pharmaceutical research process

including therapeutics. RNAi has been used to target dominant

mutant or amplified oncogenes, translocation products, signal-

ing molecules and viral oncogenes such as bcr-abl, mutated ras,

or over expressed Bcl-2. Therapies based upon RNAi may have

a number of inherent, fundamental benefits, such as harnessing

natural pathways and the potential to target virtually any pro-

tein, i.e., no limitation to drugable proteins. In a number of

studies it could be demonstrated that it is possible to deliver

siRNA intracellularly via microbubble-enhanced focused ultra-sound [62,133].

6. Conclusions and outlook

The integration of therapeutic interventions with diagnostic

imaging, to allow for local image-guided delivery, calls for devel-

opments in equipment and agents including new therapeutics.

Focused ultrasound in combination with MRI and ultrasound imag-

ing has great potential to bring ultrasound triggered drug release

to the clinic, while employing pressure and temperature sensi-

tive delivery vehicles. The preclinical data demonstrate the specific

solutions that are emerging for local drug and gene delivery in both

oncology and cardiology.

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Acknowledgement

The assistance of Dr. Sander Langereis in preparing the figures

is gratefully acknowledged.

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