Delivery in Solid Tumor Treatment
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V Vo ol l.. 4 4 ·· N No o.. 2 2 ·· 2 20 01 13 3
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013 Page 223–254
Ultrasound-Mediated Drug/Gene Delivery in Solid
Division of Engineering Mechanics, School of Mechanical and Aerospace
Engineering, Nanyang Technological University, Singapore
Submitted July 2012. Accepted for publication December 2012.
Ultrasound is an emerging modality for drug delivery in chemotherapy. This paper reviews this novel
technology by first introducing the designs and characteristics of three classes of drug/gene vehicles,
microbubble (including nanoemulsion), liposomes, and micelles. In comparison to conventional free
drug, the targeted drug-release and delivery through vessel wall and interstitial space to cancerous
cells can be activated and enhanced under certain sonication conditions. In the acoustic field, there
are several reactions of these drug vehicles, including hyperthermia, bubble cavitation, sonoporation,
and sonodynamics, whose physical properties are illustrated for better understanding of this
approach. In vitro and in vivo results are summarized, and future directions are discussed. Altogether,
ultrasound-mediated drug/gene delivery under imaging guidance provides a promising option in
cancer treatment with enhanced agent release and site specificity and reduced toxicity.
Keywords: ultrasound, drug/gene delivery, drug vehicle, bubble cavitation, sonoporation
Cancer is currently the second leading cause of mortality in the world. According to the
Surveillance Epidemiology and End Results (SEER) Cancer Statistics Review, there
were more than 570,000 deaths due to cancer and approximately 1.6 million new cases
diagnosed in 2011 in the United States . Although half of cancer patients die as a
result of malignant metastases, lack of control of the primary tumor causes a critical
failure, especially for cancers at the cervix, colon, ovaries, pancreas, and brain.
Extensive research and substantial progress in the development of anti-cancer agents
have been made in recent decades. However, success achieved in solid tumors
(accounting for 85% of all cancers) treatment is not satisfactory. Three physiological
barriers limit the efficacy and safety of chemo- and bio-therapeutics especially for the
most promising macromolecular agents (i.e., monoclonal antibodies, cytokines,
antisense oligonucleotides, and genes): blood vessel walls (vasculature occupying
1-10% of the tumor volume); collagen-rich interstitial space accounting for a large
*Corresponding author: Yufeng Zhou, Nanyang Technological University, School of Mechanical and
Aerospace Engineering, Division of Engineering Mechanics, 50 Nanyang Avenue, Singapore, 639798.
Phone: (65) 6790-4482. Fax: (65) 6792-4062. E-mail: firstname.lastname@example.org.
volume of tumors; and the cancer cell membrane, due to the heterogeneity of antigen
and receptor expression for affinity-targeted delivery of drugs . The typical diffusion
time required for macromolecules to cross a distance of 200 µm (average distance
between tumor capillaries) is longer than their average half-life time . Uneven and
slowed blood flow within tumors because of abnormal (aberrant branching and
tortuosity), heterogeneous, and inefficient distribution of a vasculature , high
interstitial pressures due to the tumor vessels’ leakage, the absence of a functional
lymphatic system , and fibrillar collagen in the extracellular matrix  further
complicate effective and uniform delivery of high-molecular-weight (molecular weight,
MW > 2000 Da) drugs. The dose of chemo- and radio-therapeutic agents for clinical
therapy usually destroys normal cells/tissues, and result in a variety of undesirable side
effects, including cardiotoxicity, immune suppression, and nephrotoxicity [6, 7].
In order to overcome this stalemate, various approaches have been explored to
maximize drug localization to the tumor while minimizing systemic toxicity .
Lately, controlling drug deposition and release in the region-of-interest by an external
signal, such as light, neutron beam, magnetic field, or mechanical energy, is gaining
attention . Ultrasound has been applied in medical diagnosis since World War II. In
recent years, it has taken on a new life in therapeutic applications [10, 11], such as
targeting or controlling drug release [12, 13]. In 1976, the cytotoxic effect of nitrogen
mustard on mouse leukemia L1210 after sonication without any mechanical damage to
cells was first observed, which cannot be explained merely by ultrasound-induced
hyperthermia . This approach is highly attractive because of wide acceptance of
ultrasound in medicine, easy and accurate focusing to the deeply seated organs in the
body, high efficiency in perturbing cell membranes and increasing their permeability,
noninvasiveness, non-viral nature, low cost, and non-ionization for theoretically
unlimited treatment . Meanwhile, there is no effect on the non-sonicated region.
Overall, this strategy provides an option in treating the localized tumors that may be
inoperative by surgery either physiologically or cosmetically.
In this paper, the progress and characteristics of ultrasound-mediated drug delivery
in the solid tumor treatment are reviewed. First, the design of an acoustically-active
drug, such as microbubble, nanoemulsion, liposome, and micelle, is discussed. Then
the underlying mechanisms of ultrasound-induced enhancement of drug delivery are
explained. In vitro and in vivo outcomes of this novel technology are summarized for
future translation into clinics. Finally, technical challenges and the potential direction
of development are discussed.
2. ACOUSTICALLY-ACTIVE DRUG VEHICLES
The most popular vehicles activated by sonication include microbubbles,
nanoemulsions, liposomes, and polymeric micelles, which have a promising
enhancement of localized delivery in solid tumors up to 5 to 10-fold over traditional
methods of delivery .
Ultrasound contrast agents (UCAs), biocompatible microbubbles less than 10 µm in
size in order to exit the heart through the pulmonary capillaries, are popular in perfusion
monitoring to measure vascular density and microvascular flow rate . Atypical dose
224 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
of UCAs for an echocardiographic evaluation is 109–1010microbubbles in a 1-2 mL
bolus intravenous injection . Microbubbles have a gas core (i.e., perfluoropentane,
sulfur hexafluoride, nitrogen) and a highly cohesive and insoluble shell (i.e., protein,
phospholipid, polymer), permitting prolonged circulation, and preventing nonspecific
removal from the circulation by the reticuloendothelial system (RES). The gas core not
only has a significant acoustic impedance mismatch for strong echogenicity, but also is
compressible for bubble cavitation (both stable and inertial cavitation, depending upon
whether the acceleration terms are dominated by the pressure of the gas or the inertia
of the inrushing liquid). For stable cavitation (SC), a bubble oscillates nonlinearly
around its equilibrium radius. The stable pulsation of fluid produces microstreaming
and shear stress, which are shown to induce transient compromise of cell membranes.
However, in inertial cavitation (IC), a bubble is reduced to a minute fraction of its
original size, and the gas within dissipates into the surrounding liquid via a violent
mechanism. UCAs, regardless if being co-administered or encapsulated with
pharmaceutical agents, can be intentionally ruptured by ultrasonic waves at a
moderately high acoustic pressure at the target sites [17, 18].
2.1.1. Drug-loaded Microbubbles
There are various ways of entrapping drugs within a microbubble (Figure 1) [20, 21].
Drugs may be incorporated into the membrane or in a shell of microbubbles. Stable and
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
(a) (b) (c)
Different approaches of loading drug/DNA into microbubbles by
(a) attaching to the membrane, (b) embedding within the membrane,
(c) bounding non-covalently to the surface, (d) enclosing inside, and
(e) incorporating into an oily film surrounded by a stabilizing layer with
a ligand for targeting (used with permission ).
strong deposition of the charged drugs in or onto the microbubbles’shell is realized by
electrostatic interactions. However, Kupfer cells, leukocytes, and macrophages have the
tendency to capture charged microbubbles, which could substantially decrease their
half-life. In addition, drugs can also be embedded into microbubbles. The advantages
of drug-loaded microbubbles are that the loaded-drugs can be released at the occurrence
of IC into the tissue due to the induced shock wave, microstreaming and microjet.
Meanwhile, the process could be tracked with sonography as the drug carriers are
essentially UCAs. However, the monolayer lipid shell (a few nm) limits loading the
hydrophobic pharmaceuticals, and may have a premature release. Although a thicker
triglyceride lipid shell and dissolved drug is a solution, it is only available for
hydrophobic drugs (i.e., paclitaxel). Biocompatible and biodegradable polymeric (i.e.,
pLGA) microbubbles possess a thick and rigid shell for a prolonged circulation half-
life, and permit a much higher loading capacity of both hydrophobic and hydrophilic
drugs whose release rate depends on the drug properties (i.e., lipophilicity and water
The aforementioned methods are most applicable to highly active drugs, such as
gene-based drugs, assuring protection of the drug and lowering the propensity of
premature release. Therapeutic genes with several thousand pairs and MW over 1
million Da cannot cross the capillary fenestrations of blood vessels . In addition,
intravenous administration is not suitable for genetic materials because of the rapid
metabolization by serum esterases. After reaching the tissue targets, genes must
penetrate cell membranes, pass cytoplasm, enter the cells’ nuclei, and then be digested
by lysosomes within the cells. Drawbacks of incorporated naked plasmid DNA(pDNA)
and pDNA-polymer complexes released from microbubbles are (a) large microbubble
size (3-7 µm for consequent short circulation time and ineffective extravasation into the
tumor); (b) the necessity to complex the pDNA with cationic polymers to prevent
degradation during fabrication; (c) low loading efficiency of pDNA (~6700
molecules/bubble) due to the limited number of cationic lipids; and (d) prematurely
release of more than 20% of the encapsulated pDNA .
2.1.2. Targeted Microbubbles
Although acoustic radiation force could promote the attachment of microbubbles to
endothelial cells in the wave propagation path, it is more attractive to selectively adhere
microbubbles to cellular epitopes and receptors of cancer or solid tumors by one or
several specific ligands, such as antibodies, carbohydrates, and/or peptides (Figure 2)
. Monoclonal antibodies have a very high specificity and selectivity to a large range
of epitopes. In contrast, peptides are low-cost and less immunogenic because of much
smaller size (5-15 amino acids). Significant enhancements of microbubbles adhesion to
activated endothelium , rejecting tissues , neovasculature endothelium ,
lymph node-related vasculature , or activated platelets  have been reported.
Simultaneous targeting to multiple ligands could synergistically increase adhesion
strength . There are two ways of coupling ligands to the microbubble shell: covalent
(being attached to the head of phospholipids directly or via an extended polymer spacer
arm), and non-covalent by avidin-biotin bridging because of the wide availability and
excellent affinity. However, since avidin carries a strong positive charge in the
226Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
glycosylate layer, the bio-distribution of microbubbles may be altered, resulting in non-
specific adhesion and introduction of undesired immune response. In comparison,
Streptavidin may be a better bridging molecule. It is found that connecting an
intermediary spacer arm with the ligand at the end as a polyethylene glycol (PEG)
molecular tether to the lipid shell indirectly is feasible, and has high specificity and
targeting. Meanwhile, the tethered ligand could also be buried in a polymeric
“overbrush” . After sonication, the ligand is instantly revealed for targeting to avoid
the undesired non-specific binding and uptake by RES.
Altogether, microbubbles can enhance transport of drugs/genes across vessel walls
and cell membranes by the localized bubble cavitation. However, their disadvantages
are the relatively low payload, and short circulation half-life (minutes for lipid-shelled
microbubbles, although longer times for the polymer types).
Although the use of microbubbles is very attractive in drug delivery, especially for its
uniqueness of combined diagnosis and targeted therapy with a high effectiveness-to-
cost ratio, it lacks an essential prerequisite for effective extravasation into tumor and
subsequent drug targeting with a sufficient lifetime in the circulation, because the pore
size of most tumors is usually smaller (380-780 nm) . An alternative solution to
aforementioned problems is developing drug-loaded nanoparticles that could
accumulate in tumor and then expand to microbubbles in situ under sonication, such as
copolymer-stablilized echogenic perfluoropentane (PFP) nanoemulsions. The
nanoemulsions are produced from drug-loaded poly(ethylene oxide)-co-poly(L-lactide)
(PEG-PLLA) or poly(ethylene oxide-co-polycaprolactone (PEG-PCL) micelles.
Dodecafluoropentane (DDFP) droplet remains as a liquid at the body temperature
although it has a boiling temperature of 29°C at atmospheric pressure, but vaporizes
under ultrasound exposure. Acoustic droplet vaporization (ADV), transiting droplet to
bubble, is determined by a certain pressure threshold [33, 34]. The ADV threshold for
PEG-PCL stabilized droplets is lower than that required by PEG-PLLA droplets. The
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
Targeting microbubble to cancer cells by connecting the receptors on the
surface with (a) an antibody, (b) an avidin bridge, and (c) a flexible spacer
arm (used with permission ).
higher echogenicity of microbubbles enables its discrimination from droplet’s status in
sonography. To prevent excessive phase transition in vasculature, infusion or injection
through a low-gauge needle should be carried out in the application of PFP
nanoemulsions. A significant volume expansion in the complete vaporization of the
droplet and consequent shrinkage of the microbubble shell may rip off a drug to the
local cells (Figure 3). All aspects considered, this method allows the enhancement of
intracellular drug delivery with high ultrasound contrast and drug loading [31, 33, 34].
Liposomes with a typical diameter of 65-120 nm are non-toxic, biodegradable, and non-
immunogenic drug delivery vehicles for both hydrophilic and lipophilic drugs, such as
doxorubicin (DOX) and vincristine [35, 36]. Local heating causes the liposome to
change from a well-ordered gel to a less-structured liquid crystalline state by inducing
thermotropic phase transitions on phospholipids (Figure 4). As a result, the cargo could
be released with significant reduction on the systemic toxicity, and premature
degradation or inactivation . The in vivo stability and accumulation at the tumor site
can be enhanced to 50 to 100-fold compared to the free drug , and small polymeric
carriers can respond to both mechanical (i.e., fracture) and thermal (i.e., temperature
elevation) activation. PEG liposomes containing DOX have been used in the treatment
of Kaposi’s sarcoma, refractory ovarian cancer, and breast cancer . Furthermore,
liposomes are also efficient as non-viral gene carriers .
Different strategies are applied for both drug loading and release from microbubbles
and liposomes. Although the gas core occupies the greater volume of a microbubble, the
drug loading capability can be enhanced by shell construction layer-by-layer or
228 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
Schematic diagram of phase transition from nanodroplets to
microbubbles, and consequent drug release into the neighboring cells
under sonication (used with permission ).
conjugation of drug-entrapped particles. In comparison, drug-specific loading
techniques are applied to construct the core of a liposome. Drug release from
microbubbles is mainly due to the ultrasound-induced IC, whereas that from a liposome
is the consequence of the heat produced by the absorption of acoustic energy.
Meanwhile, the gas nuclei will expand and dilate the monolayer boundary or the
adjacent bilayer in the strong acoustic field. If the bubble expansion-induced stress is
beyond the elastic threshold of the bounding membrane, the liposome will rupture, and
then the incorporated contents will be released.
2.3.1. Thermosensitive Liposomes
The thermally responsive macromolecular carrier, poly(N-isopropylacrylamide,
NIPAAm), can be cheaply and conveniently synthesized by the approach of free radical
polymerization . However, precisely controlled release of chemo- and radio-
therapeutic agents conjugated to its structure is difficult. Artificial Elastin-like
polypeptides (ELPs) are genetically encodable, and polypeptide-based polymers that
are recombinantly synthesized in E. Coli by over-expression of a synthetic gene,
allowing control over the sequence, chain length, and the number and location of
reactive side chains (e.g., lysine, cysteine) on the polypeptide . ELPs exhibit lower
critical solution temperature (LCST) than poly(NIPAAm). Traditional thermo-sensitive
liposomes (TSLs) are triggered at 42-45°C for drug release over ~30 min. In
comparison, the payload from low temperature–sensitive liposomes (LTSL) could be
released in the temperature ranges of 39-40°C, as shown in Figure 4.
2.3.2. Non-thermal Liposomes
Acoustically active lipospheres (AALS) do not rupture as easily as UCAs because the
thicker shell retards the wall velocity in the bubble cavitation. A targeting ligand is
necessarily attached to the liposomes to prevent them from being flushed away by the
circulating blood. However, a lower paclitaxel loading efficiency (4 mg/mL of
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
Temperature-sensitive liposomes (left) below and (right) above their
phase transition temperature, which is typically chosen near 40°C for
temperature-sensitive vehicles. Typical liposome diameter is 65-120 nm
(used with permission ).
microbubbles) may be an unfavorable side effect of the ligand attachment. Loading and
release of hydrophobic drugs in AALS could be improved by varying the amount of oil
and perfluorocarbon gas. In summary, it remains rather difficult to obtain an efficient
drug release from AALS and to also incorporate hydrophilic drugs . Meanwhile,
air-filled echogenic liposomes (ELIPs) with an average size of ~800 nm were
developed to incorporate hydrophilic compounds inside lipsomes (Figure 5). Drug
release from ELIPs is expected to occur relatively easily. When the expansion of ELIPs
becomes so high that the elastic limit of the lipid shell is exceeded, drugs can leak out.
The incorporation of short-chain lipids in the ELIPs shell results in a prolonged shell
opening and a better drug release . Because of the flexibility of the lipid monolayer,
the ELIPs shell can reseal. Therefore, successive expansion cycles can be used to
230 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
Schematics of (a) an acoustically active liposphere (AALS)
encapsulating hydrophobic drugs and hydrophobic gas; and (b), (c)
echogenic liposomes (ELIPs) encapsulating hydrophilic drugs and
hydrophobic gas (used with permission ).
improve drug release from ELIPs. AALS is conceptually similar to microbubbles,
except the smaller size for easy extravasation, relatively low encapsulation efficiency,
and limited drug release from the shell collapse because a thicker shell retards the wall
velocity associated with IC.
2.3.3. Liposomes Loaded Microbubble
A more advanced drug-carrying microbubble was achieved by attaching both
hydrophilic and hydrophobic drug-loaded nanoparticles to the structure via
avidin–biotin interactions (Figure 6) [43-45]. As ~105liposomes are bound to each
microbubble, the drug loading capacity of microbubbles can be significantly enhanced.
Attachment of polystyrene nanoparticles/liposomes onto the microbubble shell did not
hinder IC. Cells close to liposome-loaded microbubbles contained liposome fragments
after sonication. This concept can also be used to load siRNA and gene-containing
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
Schematic representation of a nanoparticles-loaded microbubble.
PEGylated nanoparticles are attached to the microbubble with an avidin-
biotin binding. Ultrasound irradiation results in the release of intact
nanoparticles (used with permission ). DPPC: 1,2-Dipalmitoyl-
sn-glycero-3-phosphorylcholine. DSPE: 1,2-Distearoyl-sn-glycero-3-
liposomes (lipoplexes) to improve their transfection efficiency. These vehicles combine
the advantages of liposomes and microbubbles and overcome individual limitations.
A micelle is an aggregate of surfactant molecules dispersed in a liquid colloid with the
hydrophilic “head” regions in contact with surrounding aqueous solution and the
hydrophobic single-tail regions in the micelle center. They are small enough to avoid
renal excretion, but permit extravasation at the tumor site via the enhanced penetration
and retention (EPR) effect . The structure of hydrophobic cores as drug reservoirs
and hydrophilic shells ensure micelle solubility in the aqueous medium but prevent
micelle aggregation. The PEG blocks with a typical length of 1-15 kDa prevent micelle
opsonization, in order to avoid micelle recognition by RES. The thermodynamic and
kinetic stability of micelle is determined by molecular interactions and the length of
hydrophobic blocks and the micelle core type, respectively. Soluble polymeric carriers
that undergo an LCST phase transition would enable targeting to heated tumors, and
hydrophobically collapse and aggregate at temperatures higher than their LCST .
The self-assembly of amphiphilic block copolymers is activated thermodynamically
and is reversible when the concentration reaches the critical micelle concentration
(CMC). However, copolymer molecules are individual (unimers) at low concentrations.
The encapsulated drug will be released from the prematurely destroyed micelles. The
CMC is primarily controlled by the length of a hydrophobic block (the greater the
length, the lower the CMC), but less sensitive to the length of a hydrophilic block.
However, an excessive concentration of the copolymer will initiate micelle aggregation
and precipitation . Micelles, whose glass transition temperature is higher than the
physiological temperature (i.e., 37°C), may survive for many hours or even days at
concentration below the CMC. In contrast, it takes only a few minutes to dissociate
micelles with a “soft” core (i.e., Pluronics). The CMC of this type of micelles is usually
low because of direct mixing of the hydrophobic drugs with the soft core. The most
outstanding attractions of micelles are the self-assembly of amphiphilic block
copolymer molecules and drug encapsulation by simply mixing rather than conjugation.
Bubble cavitation can occur, but only inside the micelle cores with an increased
threshold . In contrast, drug loading into micelles with solid cores requires a
number of more complex techniques [50, 51].
3. DRUG/GENE DELIVERYAPPROACH
The mechanisms for the ultrasound-mediated drug/gene delivery include both thermal
and mechanical effects [52, 53]. First, acoustic radiation force can help agents penetrate
through the vessel wall to tissue. Second, ultrasound may affect the morphology and
properties of the cell membrane due to bubble cavitation for drug permeation and
absorption (sonoporation). Third, ultrasound-induced hyperthermia has a significant
biological effect on cell activities and drug uptake. Last, ultrasound can alter the
performance of a drug, for example, activating light-sensitive materials of
hematoporphyrins to kill cancer cells and inhibit restenosis . The efficiency of drug
232 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
delivery is mainly determined by the acoustic intensity. However, with a higher power,
the propensity of cell lysis is increased. Furthermore, cell membrane permeability and
drug cytotoxicities can be enhanced at elevated temperatures .
3.1. Acoustic Radiation Force
Acoustic radiation force is associated with the propagation of acoustic waves through a
dissipative medium, originating from a transfer of momentum from the wave (either
from absorption or reflection) to the medium. Radiation force is proportional to the
acoustic absorption and the rate of energy being applied, and is inversely proportional
to the speed of sound in the medium. It could push drug-carrying nanoparticles toward
a vessel’s wall prior to fragmentation, and facilitate adhesion of them with receptor-
ligands. Furthermore, the non-uniform displacement of tissue in the focal zone and
transverse waves generated at the interface of tissue and fluid create a steep gradient in
shear forces, where the resulting strain works on the cell-to-cell junctions and cellular
interfaces. The penetration of various masses (small molecules, DNA, and
nanoparticles) from the fluid into the adjacent epithelium could be increased both
systemically and locally under sonication, as well as their rate of effective diffusion
through the tissues. The increase in the intercellular gaps for enhanced local interstitial
transport and reduction in interstitial fluid pressure in the core of tumors lead to
improved extravasation of large molecules and consequent anti-tumor effects, which
are correlated with more wide-spread induction of necrosis and apoptosis [17, 56, 57].
Hyperthermia induced by the absorption and dissipation of acoustic energy by tissue at
the temperature of 40–45°C has a remarkable biological effect at the subcellular and
cellular levels: decreased DNA synthesis, altered protein synthesis (i.e., heat shock
proteins), disruption of the microtubule organizing center, varied expression of
receptors and binding of growth factors, and changes in cell morphology and
attachment [58, 59]. Importantly, hyperthermia can also increase tumor blood flow and
vascular permeability, which leads to its implementation for solid tumor therapy in
combination with thermo-sensitive drugs. Even non-thermo sensitive polymeric
carriers and drugs exhibit increased localization in heated tumors because of these
physiological effects. Furthermore, when combined with chemo- and radio-therapy,
hyperthermia produces more tumor cytotoxicity.
3.3. Bubble Cavitation
At the very low acoustic pressure [mechanical index (MI) <~0.1], the microbubble
oscillates symmetrically and linearly. At MI of 0.1-0.3, the microbubble has more
expansion than compression (nonlinear oscillations) while the shell still maintains its
integrity, a state known as stable cavitation. Higher acoustic pressure of MI (0.3–0.6)
forces microbubble oscillation for eventual destruction by either gas diffusion through
compression or large shell defects through IC (Figure 7) [9, 61, 62]. However, the
specific thresholds of different types of bubble oscillations vary a lot, depending on the
bubble structure (i.e., initial size, wall thickness, and composition) and environmental
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
conditions (i.e., ambient pressure and temperature). The microbubble wall reaches a
velocity of several 100 to 1,000 m/s at bubble collapse. Nano- or micro-particles, such
as polymer particles or UCAs, can lower the cavitation threshold although inertial
cavitation can also occur in most normal tissues. However, polymeric microbubbles
illustrate different cavitation characteristics. At the low acoustic pressure, the polymeric
microbubble will not oscillate actively because of the stiff shell. If acoustic pressure is
beyond a critical value, defects or cracks will form in the shell for gas diffusion and
microbubble fragmentation, although complete destruction is rare [63, 64].
Strong physical, chemical, and biological effects induced by both types of cavitation
(SC and IC) are the mechanisms of microbubble-mediated drug delivery (Figure 8)
[65–67]. Although shock waves, microstreaming, and microjet produced in the IC are
short in time, the localized large pressure (i.e., 1000 atm) can generate transient and non-
lethal micropores in the endothelial cell, plasma, and nuclear membrane, making them
more permeable by drugs . During the bubble collapse, the local temperature can be
as high as 5000 K (close to the surface temperature of the sun) and induce chemical
changes in the medium (sonochemistry) . The most significant is the generation of
highly reactive oxygen species (ROS), such as free radicals, that can induce chemical
transformations to permeabilize the cell membrane without producing pores on the cell
membrane and affecting the cell viability. Meanwhile, endocytosis and phagocytosis in
the uptake and fusions of lipid-shelled microbubbles are important in the intracellular
drug absorption .
Blood has a higher viscosity than water, and consequently, a higher cavitation
threshold (i.e., about 1.5 to 2 times of that in water at the frequency of 2.5 MHz). If
nanoparticles (i.e., Polystyrene) are present in tumor, the cavitation threshold could be
significantly lowered. Smaller particles have more effects on such a reduction in the
cavitation threshold at the same concentration [22, 70]. However, polystyrene beads
have a virtually indefinite lifetime in the body and are not FDA approved.
234 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
(c) (d) (e)(f)
Optical frame images (a-g) and streak image (h, diameter of the bubble
as a function of time) illustrating the oscillation and fragmentation of a
microbubble. The bubble has an initial diameter of 3 mm (used with
Efficient and well-tolerated internalization of DNA and macromolecules with a diameter
no more than 37 nm is allowed by the induction of transient and reversible holes in the cell
membrane (sonoporation). The cell structure affects the sonoporation effect and the cell
survivability, although the maximum size of molecules for entry as well as the opening
period of the pores is not well known. Immediately after sonication, the MAT B III cells
showed morphological changes (i.e., smoother surface and smaller size) because of the
removal of macromolecules (i.e., glycoproteins or cell-surface receptor CD19) on the
surface . For the cells fixed during sonication, some large pores (on the order of
100 nm) but not small ones (1–10 nm) can be produced, which may be due to the abundant
microvilli on the cell surface, and the short duration of pore opening (Figure 9a). In
comparison, fresh red blood cells have clear pores on their smooth surface after sonication
(Figure 9b). Therefore, the pore size in sonoporation depends on both the ultrasound
parameters and cell types. Small pores are more abundant on the cell membrane,
facilitating the internalization of molecules whose amount is inversely proportional to their
size, but reseals more quickly after the sonication. Large pores on the cell surface may not
reseal, and ~15 s disruption on a plasma membrane in smaller cells (i.e., fibroblasts and
endothelial cells) is generally lethal. However, the nuclear membrane seems not affected
by sonoporation, but nuclear complexes only allow the passive diffusion of small particles
up to 9 nm [72–74]. Unfortunately, the leakage of cytosolic fluid could cause partial escape
of internalized molecules. IC dominates in the ultrasound-mediated DNA transfection
through sonoporation so that measuring the IC dose may be helpful in monitoring the
delivery effect [75, 76].
3.5. Sonodynamic Therapy
Sonodynamic therapy (SDT), which uses a novel sonosensitive agent derived from
chlorophyll that was originally used as light-activated chemicals for cancer therapy, is
a promising cancer treatment modality [77, 78]. These sonosensitive drugs
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
Schematic diagram of biological effects and drug release induced during
bubble cavitation through the defective tumor microvasculature.
Ultrasound can trigger drug release/activation locally by heat, light,
shock wave, microstreaming, and microjet.
(i.e., protoporphyrin IX), upon non-thermal sonication, create a short-lived cytotoxic
ROS (i.e., superoxide radicals and singlet oxygen) to induce peroxidation of membrane
lipids and irreversible alteration of the target tissue. It is assumed that sonication may
introduce electronic excitation of porphyrins and a photochemical process in the
cavitation process. The agents themselves have no antitumor ability but exhibit it only
by the sonochemistry. Therefore, much less risk of adverse effects is expected on
In summary, ultrasound directly alters the biological or physiological properties of
tissues facilitating transport and the bioavailability or efficacy of the drug or vehicle. It
also produces changes in the surrounding tissue by indirect effects acting on the vehicle.
4. IN VITRO AND IN VIVO RESULTS
Early drug delivery studies involved the co-administration of a drug and microbubbles
followed by sonication. Drug-loaded vehicles were then investigated for local release
and cell uptake simultaneously only in the sonication area. Studies demonstrated that 20
kHz ultrasound irradiation in combination with the nanoparticle and chemotherapeutic
agent 5-fluorouracil (FU) injections significantly suppressed tumor growth, and even
complete elimination 2 months later, while the control volume increased and the tumor
regrew from the peripheral points of the irradiated tumors (p < 0.05). Histological studies
of H&E stained sections of control tumor revealed viable tumor cells with well-defined
nuclei. Treatment with 5-FU alone produced small necrotic regions in the tumor, most
commonly near blood vessels. Ultrasound irradiation, in combination with 5-FU and
polystyrene nanoparticles injection (100–280 nm in diameter and concentration up to
0.2% w/w), resulted in dramatic tissue necrosis that was noted throughout the
236 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
Representative pores at (a) MAT B III and (b) red blood cells after
sonication (f = 2.25 MHz, p-= 570 kPa) in the presence of UCA (25
particles/cell for MAT B III and 1.2 particles/cell for red blood cells) by
scanning electron microscope at a magnification of 10,000 (used with
Optimization of ultrasound frequency, intensity, duty cycle, time of irradiation, and
concentration of Optison®led to 73.5 ± 3.3%, 72.7 ± 0.9%, and 62.7 ± 2.1% delivery of
10 kDa, 70 kDa, and 2000 kDa macromolecules in the MCF7 cell line, respectively, and
36.7 ± 4.9% of cell transfection, while dead cell count was only 13.5 ± 1.6% . These
results suggest that optimized treatment parameters provide efficient drug and gene
delivery to cancer cells and could be utilized in further in vivo experiments (Figure 10).
Bubble cavitation-induced convection can transport drugs or particles over several
10 µm from the vessel wall. The change in capillary permeation is strongly dependent
on cavitation dose. Acoustic pressure of 0.75 MPa at 1 MHz can produce capillary
rupture in the intact rat muscle microcirculation with high microbubble concentration,
and the pressure required increases proportionally to the driving frequency. Combining
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
Confocal microscopy of MCF7 cells loaded with FITC-dextran by
sonication (3 MHz, 20% duty cycle, 3 W/cm2, 1 minute with 200 µL of
Optison per 1 mL sample). (a) Red staining (propidium iodide) marks
nuclei of dead or dying cells, while green indicates distribution of 70 kDa
FITC-dextran. Single cells loaded with (b) 2000 kDa, (c) 70 kDa, and (d)
10 kDa FITC dextran. Bars indicate 10 mm (used with permission ).
radiation force with destructive ultrasound pulses has increased deposition of oil on cell
membranes in vitro by 10 folds in comparison to ultrasound alone .
Hemolysis, microvascular leakages, capillary ruptures, petechial hemorrhages, mild
elevations of troponin-T in blood, cardiomyocyte injury, inflammatory cell infiltrations,
and premature ventricular contractions were also observed in both in vitro and in vivo
experiments after the sonication with UCAs . The extent of these bioeffects is
influenced by several factors, such as the concentration of UCAs, the drug-delivery
method (intra-arterial vs. intravenous), the characteristics of acoustic field (i.e.,
frequency, pressure, beam size, duty cycle, exposure time), properties of the targeted
tissue, and the ultrasound imaging mode (intermittent).
Gadolinium-tagged nanoemulsion was found to accumulate in rabbit VX2 tumors for at
least six hours after intravenous injection in T1-weighted MRI and change to
microbubbles using short HIFU pulses in B-mode ultrasound image. As a result, HIFU-
induced heating in the tumors are enhanced . An ovarian carcinoma tumor
disappeared completely after four sonications with systemic injections of the
nanodroplet-encapsulated PTX, nbGEN (20 mg/kg as PTX), in two weeks. PTX alone
is tightly retained by the nanodroplet carrier in vivo, which avoids the unintentional
release to normal tissues. Thus, the unsonicated tumor grew at the same rate as control.
A similar phenomenon was also found in the pancreatic cancer model . It is
important to note that the number of metastatic sites was substantially lower in the
treated groups, indicating the potential ultrasound-induced immune response, whose
mechanism is unclear and under investigation .
Liposomes have a long half-life time in the circulation, resulting in a high-resolution
positron emission tomography (PET) image of the vascular structure (Figure 11). The
accumulation of liposomes with a vascular targeting ligand on the shell at the heart
endothelium can occur quickly (i.e., within 100 s in mouse heart). In comparison, lipid-
shelled microbubbles can circulate for a few minutes with the shell accumulating in the
liver and spleen, and the gas core exhaled through the lungs .
Acoustically active liposomes can have drug-encapsulation efficiency as high as
15%, and drug is released by 1 MHz ultrasound at 2 W/cm2for 10 s. The sensitivity
of liposomes to ultrasound stimulation could be increased by the inclusion of 4%
diheptanolyphosphatidylcholine (DHPC), which has a high correlation with the loss
of gaseous core in the acoustic field. Overall, the current encapsulation and triggered
release techniques in liposomes have a high efficiency for great potential in drug
Initially, there is no release of DOX from either LTSLs or NTSLs. After 2-min
incubation, the LTSLs began releasing DOX at a temperature of 39°C (~35% of
payload), more at 42°C (~50%), and approached 100% after 12 min. In contrast, the
NTSLs did not release any detectable levels of encapsulated DOX at 42°C after 12 min.
When LTSLs were exposed to pulsed high-intensity focused ultrasound (HIFU), a 3- to
4-fold increase in concentration of DOX in a more rapid manner was found in vivo .
238 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
In the first clinical trial of micellar drug in Japan, DOX was chemically conjugated to
PEG–P(Asp) chains and protected from digestion by erythrocytes. This micelle
manifested a higher DOX concentration in the solid tumor, but exhibited a significant
decrease in cardiotoxicity at the DOX dose of 100 mg/kg [48, 84]. In rats receiving
DOX-encapsulated micelles, the sonicated tumor grew less than the bilateral tumors,
despite large variation among each group, and incomplete tumor regression during the
time investigated. In comparison, the tumor volume increased exponentially over time
in the negative control group. The slowed tumor growth was due to either increased
DOX concentration in the vicinity of the tumor or interaction among the ultrasound, the
tumor, and DOX leached from the micelle. Ultrasound promotes the extravasation into
the tumor capillaries of any Plurogel drug carriers. The increased permeability of
angiogenic vessels and the quantity of DOX from stable micelles improved hepatic
colorectal metastases in a mouse model. If defects occur in the cell membrane,
ultrasound could enhance the drug uptake to the tumor, irrespective of whether being
released from a nearby Plurogel by sonication or having previously diffused out of the
Plurogel. Because extensive cell lysis can cause a massive release of lysosomal
enzymes and acute inflammation, there should be a compromise between the enhanced
drug uptake and cell lysis for the optimal outcome.
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
Positron emission tomography (PET) 90-min maximum intensity
projection images of (a) long-circulating liposomes, (b) short heart-
targeted peptide (Cys-Arg-Pro-Pro-Arg) coated liposomes, and (c)
microbubbles in a rat model. Particles were radiolabeled by incorporating
[18F]FDP into the particle shell (used with permission ).
Drug release from micelles could be due to either the diffusion out of micelles or
micelle perturbation/degradation by ultrasound. The released drug was quickly re-
encapsulated in the interval time of sonication pulses so as to minimize the side effects
to normal tissues (Figure 12). At pulse duration longer than the threshold, up to 10%
release of DOX is close to that in continuous wave (CW) mode . Pulsed wave is
preferred to CW by virtue of minimizing potential overheating in the clinical practice.
Longer pulses and shorter intervals could keep a high concentration of released drug
consistently throughout the sonication. The drug release from Pluronic micelles, which
has comparable characteristic times of the intracellular drug uptake, increases with the
pulse duration, but not the pulse interval. The drug re-encapsulation rate does not
depend on the pulse duration. If intervals are long enough, complete re-encapsulation
of the released drug occurs.
Optimal stability, low toxicity, and long half-life time of encapsulated DOX were
achieved by mixing Pluronic P-105 and polyethylene oxide (PEO)-diacylphospholipid.
In addition to the increased intracellular uptake, a uniform distribution of the micelles
240 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
0.5s ''on'' : 0.5s ''off''
1s ''on'' : 1s ''off''
1s ''on'' : 3s ''off''
0 20 40 6080
100 120140160 180
Fluorescence intensity, Arb, units
Profile of doxorubicin (DOX) release profiles from 10% Pluronic
micelles under continuous wave (CW) or pulsed 20 kHz ultrasound at an
acoustic intensity of 0.058 W/cm2. The decrease of DOX fluorescence
intensity was used to indicate the concentration of DOX in the aqueous
medium (used with permission ).
and drug in the tumor volume were found after the sonication via the EPR effect,
without molecular targeting (Figure 13). For stabilized Pluronic micelles, the optimal
sonication was 4–8 h after the injection . Despite excellent performance in vitro,
Plurogels induced significant cell dehydration in vivo.
Multi-drug resistance (MDR) impedes chemotherapy for metastatic diseases. One
of the well-understood mechanisms of MDR is that over-expression of P-glycoprotein
(P-gp) enables tumor cells to expel many structurally and functionally unrelated
hydrophobic anticancer drugs using the energy of ATP hydrolysis. Most of the P-gp
modulating agents, such as verapamil, are cytotoxic, and lack specificity due to the
abundance of P-gp. The enhancement of intracellular uptake and cytotoxicity of DOX
as well as inhibition of drug efflux produced by ultrasound hyperthermia was far
better than that produced by the P-gp modulator. Simultaneously, Pluronic micelles
also shielded the drug uptake. These two processes compete with each other,
depending on micelle concentration and micelle-drug interaction. Pluronic F-128 at
concentration up to 1 wt.% has obtained clinical approval from FDA. Comparative
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
Subcellular trafficking of the Ruboxil (Rb) (a) did not penetrate into the
cell nuclei (indicated by arrow), and (b) in 1% Pluronic P-105 micelles
effectively accumulated in the cell nuclei of the multi-drug resistant
ovarian carcinoma cells (used with permission ).
studies of chemotherapy with and without the use of ultrasound (20 and 70 kHz) have
been carried out in mice bearing highly MDR human colon KM20 tumors .
Sonication alone does not produce tumor regression or affect the tumor growth rate
significantly. Without anti-cancer drug, sonication in combination with the
nanoparticles resulted in a temporal decrease of tumor volume. However, the tumor
regression occurred several days later, suggesting that it was due to mechanical
damage to tumor blood vessels by bubble cavitation (SC and IC). There was a
statistical difference between sonicated and control tumors (p < 0.05) for up to 2
months in mice. No metastatic tumors were observed in the internal organs.
Drug release from micelles decreases with driving frequency (i.e., from 20 kHz to
3 MHz) and increases linearly with acoustic intensity for the stabilized Pluronic
micelles. However, there is no acoustic intensity threshold in drug release from
micelles, suggesting that IC is not the sole mechanism involved. Decreased drug efflux
and increased nuclear transport resulted in a dramatic sensitization of the MDR cells in
the presence of Pluronic micelles, more than that on non-cancerous cells. Despite
decreased intracellular drug uptake by Pluronic micelles due to the induced deactivation
of drug efflux pumps by the decreased energy in the MDR cells, the high-growth
inhibition was almost the same for micellar-encapsulated and free drug, possibly
because of the cytostatic action of Pluronic micelles overcoming the cytotoxic action of
the drug. Such an inhibition effect on MDR cell growth was more significant than that
on drug-sensitive cells. The presence of Pluronic unimers rather than that of micelles at
the absence of drugs could enhance both cytostatic and cytotoxic actions. All in all,
localized and controlled drug delivery to drug-sensitive and MDR tumors by the
synergy of drug, micelle, and ultrasound techniques that can effectively deliver drugs is
feasible, and will be further evaluated in the coming in vivo experiments.
4.5. Gene Expression
Microbubbles (i.e., Levovist®, Albunex®, and Optison®) with ultrasound can have up
to 6-fold increase in the beta-Gal gene transfection to vascular smooth muscle cells
(VSMCs) in comparison to the mixture of genes and VSMCs. If adenovirus genes are
loaded on the lipid shell of microbubbles, a 10-fold increase in a mouse heart model
can be achieved. Delivery to a specific site can be aided using the targeted
microbubbles. Bubble cavitation (SC and IC) leads to a fusion between viral vectors
and cell membranes rather than a forced deposition of intact virus particles inside the
cytoplasm. In this way, the virus can unpack and easily proceed to the nucleus. Also
remarkable was the rather uniform gene transfer in the cardiac tissue, probably due to
the ability of adeno-associated virus (AAV) serotypes to cross blood vessel barriers
and spread through the extracellular matrix . Moreover, adverse immune responses
can be drastically reduced, as the virus is shielded from the immune system. A
transfection efficiency of 50% was comparable to that achieved using lipofection. As
a physical method, ultrasound does not have limitations in plasmid DNA (pDNA)
uptake into cells by lipofection, such as the net charge of the cationic lipids, DNA
complexes and concentrations, the pH, and concentration of electrolytes. The
optimization of different cationic lipids used for each lipofection experiment leads to
242 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
inconsistent results . Although relatively large amounts of pDNA are needed for
competitive transfection rate in sonication, the in vivo site specificity generally
outweighs this small disadvantage .
It was observed that macromolecules with different MWs have similar penetration
capability through plasma membrane . In MCF7 cells, the overall difference
between the delivery rates of 10 kDa and 2000 kDa dextrans was about 20%. However,
confocal microscopy revealed that intracellular distribution of the delivered
macromolecules depended upon their size. Macromolecules of 2000 kDa did not
penetrate through the nuclear membrane, while 10 kDa macromolecules were
uniformly distributed within the cells. Macromolecules of MW 70 kDa can penetrate
through the nuclear membrane; however, nuclear fluorescence of FITC-dextran is
weaker than cytoplasmic, suggesting that 70 kDa is close to a threshold of nucleus
penetration. At the same sonication conditions, SK-BR-3 human breast carcinoma cells
demonstrated only 3.5% of dead rate, while the value for HeLa human cervix epithelial
adenocarcinoma cells was about 40% . Moreover, KM20 human colorectal
carcinoma cells demonstrated more than 70% loading with macromolecular drugs, and
only 5% of dead cells, but 10% cells expressing EGFP24 h after sonication . Shock
waves have been shown to increase cellular uptake of both small (i.e., adriamycin and
fluorescein) and macromolecules (i.e., fluorescein-labeled dextrans, ribosome-
inactivating proteins gelonin, and saporin). Utilizing 1 MHz ultrasound at a spatial
average peak positive pressure of 0.41 MPa (~10 W/cm2), a better transfection of 50%
in the living cells can be achieved, similar to the outcome of lipofection. Therefore,
ultrasound-induced gene delivery may be more or less suitable for different cell types
due to the natural variation, and further optimization of exposure conditions may
4.6. Sonodynamic Therapy
SDT can enhance the toxicity of certain compounds (i.e., haematoporphyrin,
pheophorbide, erythrosin B, and ATX-70). For example, 25 kHz ultrasound caused
lethal damage to leukemia cells in the presence of merocyanine 540 in vitro.
Hematoporphyrin, the most common photodynamic sensitizer, with 1.92 MHz
ultrasound at intensities of 1.27 and 3.18 W/cm2 enhanced the death rate of mouse
sarcoma and rat ascites 130 tumor cells from 30% to 99%, and from 50% to 95%,
respectively. The inhibitory effects of hematoporphyrin derivatives on the tumor cell
growth were confirmed by the cell’s morphological changes, cytochrome C oxidase
activity, and degradation of DNA. No cytotoxicity and structural modification of
hematoporphyrin derivative in sonication was observed on in vitro human colorectal
adenocarcinoma cells (HT-29) and Chinese hamster ovary cells. Sixty seconds of low
frequency ultrasound (270 kHz) at intensities of 0.15, 0.3, and 0.45 W/cm2and
Photofrin II (commercial and purified hematoporphyrin derivative) decreased the
survival of HL-60 cells, 92.9 ± 1.5% vs. 49.6 ± 5.1% (without vs. with Photofrin),
82.3 ± 2.2% vs. 34.5 ± 3.1%, and 77.0 ± 7.2% vs. 27.4 ± 3.0%, respectively. In an SDT
experiment (450 kHz at an intensity of 0.3–0.5 W/cm2), the survival rate of MT-2 cells
was inversely proportional to the amount of Photofrin. The corresponding values of
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
peripheral mononuclear cells in the blood of acute-type adult T cell leukemia patients
after ultrasound exposure (0.3 W/cm2, 60 s) alone, and together with 100 µg/ml of
Photofrin were 69.4 ± 22.5% and 30.0 ± 23.0%, respectively. In contrast, there was no
significant difference of survival rates of normal human peripheral mononuclear cells
between ultrasound-treated groups with and without Photofrin. Therefore, SDT may be
an effective and extracorporeal blood treatment modality for acute-type adult T cell
leukemia patients .
5. DISCUSSION AND FUTURE DIRECTIONS
5.1. Clinical Translation
The ultrasound-mediated drug delivery uses the same microbubbles or liposomes for
diagnosis and therapy at the tumor-specific region. Each component of this modality
fulfills its own function: drug vehicles target to specific tumors precisely; ultrasound
triggers the drug release in a temporally controlled manner, enhances drug diffusion
through the vessel wall for a more uniform distribution in the tumor, and perturbs cell
membranes for the intracellular drug uptake. Therefore, it is progressing toward a
feasible approach for solid tumors (i.e., breast, colon, ovaries, uterus, and larynx)
treatment with the following important and attractive aspects : (a) Microbubbles or
echogenic agents have already been used in clinical sonography. (b) It is possible to
selectively target drug delivery to the sonicated area at the deeply seated tumor using
image guidance (i.e., ultrasound, computed tomography (CT), magnetic resonance
imaging (MRI), or PET) with a high positioning accuracy, which may be beneficial in
noninvasive treatment of a localized tumor. (c) Ultrasound exposure is a noninvasive
and extracorporeal method in comparison to radiofrequency and microwave ablation,
which require insertion of interstitial needle or antenna. (d) Delivery of various agents
(i.e., small molecules, DNA, and nanoparticles) to a variety of tissues with satisfactory
therapeutic efficacy by mechanical, thermal effect, or combination is possible. (e) Drug
release was rapid at the sonicated region only, and such target specificity is much better
than current chemotherapy. (f) It could also be combined with other adjuvant therapies
to avoid or reduce metastasis. However, the drawback is relatively small volumes (a
few millimeters in the lateral by about 1 centimeter in the axial directions). In clinical
treatment, the HIFU focus is required to be scanned throughout the whole target
volume, which may take hours depending on the target size and treatment planning.
Although extensive studies have been carried out both in vitro and in vivo with
promising results, little significant progress has been made in the clinical practice.
During the last decade, combined diagnostic and therapeutic ultrasound (i.e., HIFU) has
been tested in clinical trials and has attracted growing interest, and concomitant
instrumentation is available on the market, such as ExAblate (InSightec, Israel),
Achieva (Philips, Briarcliff, NY, USA), JC (Chongqing Haifu (HIFU) Technology,
Chongqing, China), and Theraclion (Malakoff, France). Ultrasound-mediated drug
delivery may be rapidly applied to humans with only changes of operation parameters
in existing HIFU devices. However, this technology is still in its infancy, and that road
ahead presents a number of exciting opportunities as well as pitfalls that will need to be
244 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
5.2. Ultrasound Parameter Optimization
Amajor challenge is to understand the mechanisms involved in this technology in order
to optimize the ultrasound parameters. For example, for effective molecular
internalization facilitated by the increased cell permeability, the plasmid should be close
to the target because the cells return almost immediately to their initially impermeable
state after sonication. Therefore, the transient pore opening limits the efficiency and
quantity of delivery. If the IC-induced pores in the cell membrane are too large to reseal
quickly, the cell may lose vital cytoplasmic compound. The primary cause for cell death
of sonoporation is likely to be a mechanical effect, resulting from shear stresses to the
cell membranes by violent IC or microstreaming, which was supported by the
observation that a large number of cells disappeared shortly after exposure
(approximately ten minutes), but before any slow biochemical processes (i.e., oncosis
and apoptosis). Therefore, there should be a tradeoff between these two phenomena for
high rate of drug penetration through the membrane into cell cytoplasm and cell
nucleus, as well as sufficient cell viability. The dominant effects need be monitored
throughout the whole process of drug delivery. Bubble cavitation (SC and IC) activities
need to be detected as the dose of the mechanical effect, and then used as feedback in
a closed-loop control system for effective and consistent outcome. Similarly, the
temperature elevation also needs to be measured when temperature-sensitive drug
vehicles are used.
5.3. Drug/Gene Vehicle Optimization
In order to improve the performance of ultrasound-mediated drug/gene delivery, efforts
are focused to (a) enhance the stability of the vesicle, (b) increase drug/gene loading,
and (c) improve targeting [41, 42]. Air sensitivity in liposome can be increased by
adding short-chain lipid (i.e., DHPC) without entrapment of air or calcein marker.
Although DHPC can destabilize bilayers by lowering their lysis tension, such a
phenomenon is not the critical factor for lytic events. A general drawback of
microbubbles as drug carriers is the rather small space available for agents loading.
Targeting ligands may prevent an efficient loading on the shell with agents, especially
via electrostatic interactions between the drug and the shell. Drug release is restricted
to the sonicated areas with low acoustic intensity, and therefore allowing a low drug
dose. Their relatively short in vivo half-life limits the therapeutic irradiation. Because
microbubbles have a broad size distribution, they will not respond evenly to a given
ultrasound frequency. Small sub-microbubbles might be able to extravasate more easily,
especially in the tumor vasculature (380-780 nm of pore size), but can be rapidly taken
up by RES [90–92]. In contrast, larger microbubbles can carry a higher drug payload
and are more easily destroyed at relatively low acoustic intensities .
An important advantage of ELPs over other thermo-sensitive carriers is that drug
accumulates because of the phase transition of the ELP rather than triggered drug
release . Therefore, a concentration gradient is not required to drive accumulation,
and ELP-drug injected at a low concentration aggregates in the heated tumor. Second,
in comparison to antibody and other affinity targeting approaches, the aggregation of
ELP-radionuclide conjugates can directly target the heated tumor microvasculature, and
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
circumvent the barriers associated with extravasation. Third, the clinical
implementation does not require any concomitant development in hyperthermia.
Finally, the aggregated ELP will resolubilize due to reversion of the phase transition
after cessation of hyperthermia.
Electroporation, a popular experimental procedure in gene transfection, requires
manipulating electric fields by setting and positioning two electrodes to cells. In
comparison, sonoporation requires a much simpler setting especially in vivo . Bubble
cavitation (SC and IC) may damage endothelial cells in both arteries and capillaries,
through which pDNAcan be transfected effectively into cardiomyocytes. Due to its small
size, siRNA could pass through the coronary arteries and arterioles, resulting in efficient
transduction from the bloodstream into arterial/arteriolar walls including the smooth
muscle cells. Although pDNA permeates endothelium slower, its transduction is
predominant at the capillary bed with intense IC. The efficacy of gene transduction
depends on the types and properties of the cells, because neuroectoderm and limb ectoderm
cells tend to express exogenous genes after a relatively short period of sonication. As part
of such differences of gene transduction efficacy, it will be important to regulate the state
of diffusion or the retention of injected solution to the target. A higher concentration of
injected DNAmay increase the viscosity of the DNA-microbubble mixture, and lead to the
failure of increment of gene transduction efficiency in neural tubes.
The poor water solubility of some sonosensitizers, mainly due to the physical-
chemical properties of hypocrellin, leads to the easy aggregation in aqueous media and
limits their clinical applications. As a result, their concentration in a specific target may
not be sufficient for the therapeutic requirement. One of the solutions is nanotechnology
in medicine, that is able to manipulate molecules and supramolecular structures to
produce programmed functions. Encapsulated sonosensitier by hydrophobic
nanoparticles has been shown to enhance the circulation time and prevent uptake by
RES. The enormous surface area of the nanoparticles can be modified for an array of
functionalities with diverse chemical or biochemical properties. In addition, presence of
nanoparticles of the appropriate size and amount in the liquid could provide nucleation
sites, decrease the cavitation threshold, and increase the liquid temperature in the
sonication because of their surface roughness. Gold nanoparticles are highly attractive
for their low toxicity, good uptake by mammalian cells, and antiangiogenetic properties.
In the presence of gold nanoparticles, the relaxation time of protoporphyrin IX
increases, favoring efficient production of ROS . Inorganic nanoparticles, such as
titanium dioxide (TiO2), have a strong interaction with light for SDT. Although TiO2
cannot be used as a SDT drug alone because of its insufficient selectivity and low
efficiency in cancer cells, it can work as an effective sonocatalyzer in the treatment of
bladder cancer and in glioma cell lines . Since the long-term side effects of
inorganic nanoparticle accumulation are still unknown, luminescent silica nanoparticles
that decompose in the aqueous medium in hours may be utilized as a sonosensitizer.
5.4. Bioeffects of the Cell Membrane
Disruption of the cell membrane is one of the mechanisms of ultrasound-mediated
drug delivery. Sonoporated cells may respond to the disrupted intracellular
equilibrium by up-regulating cytoplasmic signals related to apoptosis and cell-cycle
246 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
arrest. Ultrasound can also generate calcium transients (i.e., changes of intracellular
calcium concentration, oscillations, and spatial waves of [Ca2+]i). Ca2+is an
important messenger involved in regulating many cellular processes and functions,
such as gene transcription, cell proliferation, fertilization, metabolism, cell migration,
wound response, and phagocytosis . Bubble cavitation (SC and IC) produces an
intracellular increase of [Ca2+]i hundreds of µm away in a few seconds. [Ca2+]i
activities affect cell-cell contact, electrical resistance, ZO-1 tight junction protein
migration from intracellular sites to the plasma membrane, and tight junction
assembly in epithelial and endothelial cells . Meanwhile, mitosis of the cell cycle
is halted through down-regulation of the Cdc-2 protein. Apoptosis may be triggered
through mitochondria and promote cytochrome-c into the cytoplasm via poly-ADP-
ribose polymerase (PARP) cleavage by the activation of the caspase-3 protein.
Resealing of the cell membrane is not a simple self-reunion procedure, since the
scattered patches on the same size order as membrane pores were observed on
the resealed cells . It is hypothesized that an exocytotic patching mechanism may
be involved so that cytoplasmic vesicle will first be delivered to the sonoporation site
and then fused together with the membrane, which is similar to the [Ca2+]iinflux into
Another barrier of interest is DNA trafficking to the nucleus. Ultrasound can deliver
macromolecules into the cytosol, but does not promote their transport into the nucleus
. Microscopy illustrated that intracellular DNA was excluded from the nucleus,
which is consistent in the other gene therapy methods. Therefore, understanding the
intracellular pDNA trafficking would optimize the ultrasound-mediated gene
5.5. Other Applications
The blood-brain barrier (BBB) consists of tight junctions around all capillaries in the
central nervous system (CNS), which separates circulating blood from the brain
extracellular fluid. Because of the presence of multiple endogenous transporters, BBB
allows a selective entry of small-size nutrients and minerals across it, but restricts the
entry of large substances, such as neuropharmaceutical agents for CNS disorders (i.e.,
Alzheimer’s disease, dementia, Parkinson’s disease, mood disorder, AIDS, viral and
bacterial meningitis), for maintaining the internal milieu of the brain. Therefore, several
novel approaches, such as nanoparticles, liposomes, antibody-mediated delivery
approaches and application of genomics  are in development for brain target.
Recently, it was found that transcranial focused ultrasound can transiently permeabilize
the BBB, increasing passive diffusion in the presence of microbubbles [100-102].
Subsequent application of an external magnetic field then actively enhances
localization of a chemotherapeutic agent monitored by MRI. Combining these
techniques significantly improved the delivery of 1,3-bis(2-chloroethyl)-1-nitrosourea
(BCNU) to rodent gliomas for suppressing tumor progression without damage to
normal tissue as indicated in MRI and histology. This novel technique promises a more
effective and tolerable means of tumor therapy, with lower therapeutic doses and
concurrent clinical monitoring [103, 104]. Furthermore, it can also be used for the
treatment of thrombus, restenosis, and angiogenesis.
Journal of Healthcare Engineering · Vol. 4 · No. 2 · 2013
The use of ultrasound for the delivery of drugs to, or through, the skin is commonly
known as sonophoresis or phonophoresis. The use of high-frequency therapeutic
ultrasound (≥ 0.7 MHz) for sonophoresis (viz., high-frequency sonophoresis, HFS)
occurred as early as the 1950s, while low-frequency sonophoresis (LFS, 20-100 kHz)
has been investigated during the past two decades. Although HFS and LFS are similar
to the concept of utilizing ultrasound to increase the skin penetration, the underlying
mechanisms are different (bubble cavitation (SC and IC) and mechanical effect
dominating LFS and HFS, respectively). Aside from the benefits of avoiding the hepatic
first-pass effect, and higher patient compliance, the additional advantages include strict
control of transdermal penetration rate, rapid termination, low risk of infection, less
anxiety-provoking or painful than injection, and not immunologically sensitizing,
despite potential disadvantages of minor tingling, irritating and burning. Both HFS and
LFS improved current methods of local, regional, and systemic drug delivery or even
vaccine in the future [105, 106].
Overall, the research and application in ultrasound-mediated drug/gene delivery have
advanced tremendously during the past decade, and are expected to continue. Future
technological growth may be concentrated on, but not limited to, the following aspects:
(a) the understanding of the interaction of ultrasound with acoustically activated drug
vehicles and the subsequent bioeffects; (b) the development of drug vehicles for
effective penetration through endothelial junctions that are typically damaged in
systemic inflammation or tumor vasculature to target specific receptors rather than on
the vascular endothelium (high specificity), more effective drug/gene loading, and
sufficient circulating lifespan; (c) response of cancer/tumor cells to ultrasound
(sonoporation) and drug (apopotosis); (d) optimization of ultrasound parameters for
drug diagnosis and release in different organs with tradeoff between cell viability and
drug/gene transduction; (e) in situ real-time control by detecting the IC, temperature, or
drug concentration; and (f) large animal experiments and translation into clinical trials.
The most common cancer treatment, chemotherapy, is often limited by its cytotoxic
effects on normal tissues, particularly under high doses. Therefore, it is highly desirable
to reduce the dosage or frequency of administration by enhancing the effectiveness of
drugs to the specific target. The application of ultrasound in delivery of several
therapeutic classes (i.e., chemotherapeutic, thrombolytic, and DNA-based drugs) acting
via bubble cavitation (SC and IC), radiation forces, and/or heat has recently gained
impetus. The advantages of this novel technology include noninvasiveness, low cost,
easy operation, good focusing and penetration inside the body, and no radiation. Imaging
methods (i.e., sonography and MRI) may be particularly helpful in defining the target,
determining local drug concentration, and evaluating efficacy and temperature. Recent
successes suggest that ultrasound may be a valuable therapeutic tool for drug delivery by
lowering the administration dosage of an anti-cancer drug. However, this technology is
still in its infancy. Progression to clinical implementation will depend on the success of
resolving several key issues. The core is the appropriate design of drug vehicles and
assemblies, such as microbubble, liposomes, and micelles, capable of carrying sufficient
248 Ultrasound-Mediated Drug/Gene Delivery in Solid Tumor Treatment
payload, yet maintaining specific response to acoustic energy. Conjugating tumor- and
disease-specific antigens or binding ligands to the surface of drug vehicles can improve
their specificity for targets without compromising either echogenicity or payload
capacity. Prolonged circulation and minimal removal by RES are also required. Only by
overcoming these obstacles can ultrasound translate from bench to bedside. An
increasing body of knowledge has been acquired on the interaction between ultrasound
energy with tissues, therapeutic agents, and drug carriers for enhanced therapeutic
efficacy and efficiency. With advances in the transducer and sonography technologies, a
variety of ultrasound-based therapeutic applications have been and will continue to be
developed. For technical improvement and translation from one tissue type to another,
an in-depth understanding of the ultrasound mechanisms and physical characteristics of
the tissues, both in vitro and in vivo data, and mathematical models to optimize the
treatment protocols will be required. It must be noted that this technique has mainly been
tested in-vitro or in small animal studies to date. There is a great need for large animal
and human studies. Moreover, ultrasound-mediated gene delivery seems to outweigh
drug delivery in certain applications, such as cardiovascular disease treatment. All in all,
ultrasound-mediated drug/gene delivery is a promising technology that attracts the
interests of both scientific research and medical applications.
CONFLICT OF INTEREST
The author indicated no potential conflicts of interest.
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