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European Journal of Radiology 70 (2009) 242253
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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: marcel.bohmer@philips.com(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:marcel.bohmer@philips.comhttp://dx.doi.org/10.1016/j.ejrad.2009.01.051http://dx.doi.org/10.1016/j.ejrad.2009.01.051mailto:marcel.bohmer@philips.comhttp://www.elsevier.com/locate/ejradhttp://www.sciencedirect.com/science/journal/0720048X8/7/2019 Bhmer et al 2009
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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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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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M.R. Bhmer et al. / European Journal of Radiology 70 (2009) 242253 249
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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250 M.R. Bhmer et al. / European Journal of Radiology 70 (2009) 242253
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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M.R. Bhmer et al. / European Journal of Radiology 70 (2009) 242253 251
Acknowledgement
The assistance of Dr. Sander Langereis in preparing the figures
is gratefully acknowledged.
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