Augmentation of targeted delivery with pulsed high intensity focused ultrasound

Abstract

This paper reviews the enhanced delivery of genes, drugs and therapeutics using ultrasound. It begins with a general overview of the field and the various techniques associated with it, including sonophoresis, hyperthermia (with ultrasound), sonoporation, and microbubble assisted transvascular and targeted delivery. Particular attention is then paid to pulsed high intensity focused ultrasound drug delivery without the use of ultrasound contrast agents. Feasibility and mechanistic studies of this technique are described in some detail. Conclusions are then drawn regarding possible mechanisms of this treatment, and to contrast with the better known treatments relying on injection of ultrasound contrast agents.

• High intensity focused ultrasound
• gene therapy
• targeted drug delivery

View correction statement:

Corrigendum

Introduction

A variety of physical methods have been investigated for enhancing the delivery of therapeutic agents in vivo, including hyperthermia [1], radiofrequency ablation [2], electric fields (electroporation) [3], magnetic fields [4], increased vascular pressure [5] and even propelling plasmid-coated particles ballistically into the targeted tissue using a gene gun [6]. These methods have demonstrated various degrees of success in increasing the uptake of therapeutic agents, whether administered systemically or locally, however, they remain limited in scope and impact.

Ultrasound has long been viewed as an attractive technique for physically altering tissue permeability properties for improved uptake of therapeutic agents, including DNA. This is because it is non-invasive and has potential for accessing tissue deep in the body. It is unclear where the idea for this initially arose, but it may well have stemmed from the study of sonophoresis (also known as phonophoresis), a technique of transdermal drug delivery whereby agents are purportedly driven through the skin using ultrasound, an effect that has been known and used for many years (see Mitrigotri [7] for a review). Targeting of bulk tissue by ultrasound is a more current effort that arose in response mainly to transport problems in two tissue types: brain and solid tumor. Whereas larger molecular agents are almost completely impeded from getting into the brain by a very tight layer of endothelial cells (the blood brain barrier), in many tumors the endothelium is leaky in comparison to normal tissues, resulting in enhanced penetration and retention of molecules in the 40–100 nm range [8–10]. However, due to high interstitial fluid and oncotic pressures, such molecules fail to penetrate throughout the tumor [10–12]. Ultrasound enhanced delivery to bulk tissues has also been tried in muscle [13], the eye [14], [15], and blood clots [16], among others.

High intensity focused ultrasound (HIFU) is a technology which employs geometric and/or electronic focusing to achieve very high spatial intensities, on the order of 10²–10⁴ W/cm², over a small focal region of ten or so mm³. There are three known major consequences of having this amount of acoustic power in such a small region of tissue [17]. First, the absorption of acoustic energy results in a steep rise in temperature capable of denaturing proteins and causing coagulative necrosis of the tissue [17], [18]. Second, the transfer of momentum from the decaying wave results in a radiation force that can move tissue up to a few hundred microns, or cause the circulation of fluids [17]. Third, oscillating pressures can effect phase changes, bringing gases out of solution and causing cavitation, which, in turn, can serve to concentrate mechanical forces and disrupt tissue structures [17], [19], [20]. It is unclear whether this short list exhausts all possible interactions or not, as there has been much speculation regarding alternative mechanisms, but there exists little consensus beyond these three.

In this paper, ultrasound enhanced drug delivery techniques are reviewed, with particular emphasis on our work regarding pulsed HIFU enhancement without injection of ultrasound contrast agent (UCA) microbubbles. Recent studies of this technique are covered in some detail, for comparison with the better known and more widely accepted techniques involving deliberate introduction of UCA microbubbles. From this review, conclusions are drawn as to the possible mechanisms of pulsed HIFU enhanced drug delivery with and without injected microbubbles.

Therapeutic thermal effects of ultrasound and HIFU

The vast majority of clinical applications of HIFU have been for thermal ablation of tissue, especially tumors. Generating sufficient heat for ablation requires exposures of a few minutes, with duty cycles approaching 100%, the exact timing depending on the acoustic power being used [21]. The focused beam passes through the skin and other non-target tissues over a wide area with low acoustic intensities thus ideally avoiding tissue damage in the near field; however, at the focal point, intensities are sufficient to destroy the tissue. The advantages of HIFU over more invasive surgical procedures are numerous, including limited blood loss and infection, elimination of visible scar formation and the reduced risk of other complications [18], [19], [20], [21], [22]. HIFU ablations can also be carried out on an outpatient basis, where cost and recovery times are significantly reduced [21], [23]. The main issue is proper targeting of the treatment, as HIFU treatment too close to skin, bone, nerves, or gaseous bodies such as bowel and lung can have serious consequences [24]. Today, HIFU is being used for ablating prostate cancer [22] and uterine fibroids [23], and its use is currently undergoing clinical trials for ablating other types of cancer, including that of the breast [25], and liver [26]. HIFU exposures can be targeted to specific tissues, organs or tumors using a number of imaging modalities, including diagnostic (B-scan) ultrasound [27], X-ray computed tomography (CT) [28] and magnetic resonance imaging (MRI) [29]. HIFU exposures have also been shown to be capable of inducing hemostasis [30], [31].

Although not a focus of this review, the thermal effects of ultrasound have been investigated for drug delivery by means of delivering a local hyperthermic dose to deep tissues, either alone or in conjunction with RF fields [32]. This technique uses a low spatial average acoustic intensity, and may or may not require HIFU sources. Other crossover techniques employing ultrasound induced hyperthermia include the deployment of drugs from temperature dependent liposomes [33] and targeting via heat shock promoter Hsp70 [34].

Microbubble mediated ultrasound drug delivery

With the use of HIFU comes the possibility of very high peak spatial intensities, sufficient to produce cavitation. Heat is always produced, but may be deliberately limited by pulsing the HIFU beam to allow thermal diffusion and perfusion processes to remove the heat. Cavitation effects will then dominate the interaction. Compared to equilibrated water under normal conditions, cavitation in most bulk tissues is difficult to produce due to a lack of available dissolved gases, cavitation nuclei and space for bubbles to form and grow [19]. The resulting effects are therefore hard to predict and control, and, given the intensities involved (10³–10⁴ W/cm²), can be extremely destructive to soft tissues [19]. The use of ultrasound contrast agents in therapeutic ultrasound is a strategy designed to address these problems. These agents consist of tiny gas bubbles encapsulated in a layer of lipid or protein that travel throughout the bloodstream and provide acoustic contrast due to the strong impedance mismatch between gas and liquid. They also induce cavitation bio-effects wherever they are found, generally in the bloodstream, at much lower acoustic intensities than would otherwise be required. Many studies have made use of these agents in an effort to control cavitation for the delivery of therapeutics [35–38].

Intracellular delivery

Sonoporation is a term used to describe the direct and transient opening of the cell membrane typically associated with ultrasound applied in the presence of cavitation microbubbles [38]. If performed successfully, the pores induced in the cell membrane re-seal after a few minutes [39], leaving the cell viable but allowing sufficient time for agents to diffuse to the interior. However, it is an accepted risk that some fraction of the targeted cells will not survive treatment. Reports vary on the efficiency and safety of the technique even in the in vitro setting [40]. Nevertheless, the effect has been consistently demonstrated in vitro. In vivo work has also been performed [41], [42], most of which was geared towards gene delivery applications. Further information can be found in review articles [38], [43].

Transvascular delivery

The endothelial barrier, which controls how substances leak from the blood stream to the interstium, is a favorite target of researchers seeking to improve drug delivery, and is also an obvious candidate for microbubble–tissue interaction. This barrier is formed from a single layer of cells joined together by tight junctions. Normal transport across this layer is via fenestrations (small areas where the cell thins to allow diffusion), pores (gaps in the cell-cell junction), or active transport via vacuoles [9]. HIFU in concert with UCA microbubbles has been used extensively to disrupt this layer, particularly in the brain [44]. While the success of blood brain barrier disruption in this fashion has been rigorously demonstrated, safety is another issue. Most treatments were shown to be accompanied by some amount of micro-hemorrhage [45]. This might explain the recent push towards avoiding the inertial cavitation (imploding bubbles) regime and pursuing instead the non-inertial cavitation (oscillating bubbles) regime [45], [46]. The technical challenges of targeting ultrasound through the skull are further obstacles to overcome before this becomes clinically relevant, although this work is well underway [35], [47], [48].

An obvious extension to the use of injected microbubbles for endothelial layer disruption is to employ the same microbubbles directly as carriers [36]. The drug or DNA may be encapsulated inside the shell along with the gas, or incorporated directly into the shell itself [49]. Sometimes targeting ligands may also be incorporated, and a low level ultrasound pulse may be used to ‘push’ the microbubble to the vascular wall [50]. A strong ultrasound pulse is then used to destroy the microbubble, presumably resulting in the simultaneous opening of the endothelial barrier and the distribution of agents. While the concept would seem sound, it remains to be seen whether sufficient amounts of any therapeutic agent can be delivered in this fashion, and, with the direct use of cavitation, some level of collateral damage is impossible to avoid. Reviews of this literature are available elsewhere [36], [51].

Pulsed HIFU enhanced delivery without microbubble injection

Although the majority of groups using ultrasound for drug delivery enhancement make deliberate use of microbubbles by injecting UCAs, our group has sought to avoid the associated tissue damage by treating with ultrasound alone. The remainder of this paper will review, in some detail, the feasibility studies that were carried out in our lab over the past 5 years. Most of these studies were done in mouse models using a customized version of a commercially available HIFU system (Sonoblate 500; Focus Surgery, Indianapolis, IN). The probe consists of a therapeutic 1-MHz transducer aligned collinearly with a 10-MHz imaging transducer, each with a focal length of 4 cm, mounted on a scanner with one linear and one azimuthal degree of freedom. The therapeutic transducer is concave and spherical with a diameter of 5 cm; the aperture of the imaging transducer is 0.8 cm. A maximum acoustic power output of 120 W is available from the therapeutic transducer. The focal zone of the therapeutic transducer is ellipsoid, with a radial diameter of 1.38 mm and an axial length of 7.2 mm. Based on these dimensions, in order to treat a significant tissue region, treatment was always applied to multiple sites in a grid pattern with a 2 mm horizontal and vertical separation distance, completing a full treatment (∼100 or 120 pulses) at one raster location prior to moving to the next. The radiation force technique [52] was used (in water) to calibrate total acoustic power (TAP) levels. The overall focusing factor of the therapeutic transducer is approximately 1.3 × 10³. Mouse exposures were carried out with the anesthetized mouse suspended vertically in a tank of degassed water, typically maintained at 37°C. (Thermoneutrality in mice is achieved with an ambient temperature in the range of 29°C to 33°C in air [53]. Core body temperatures range between 36.5°C and 37°C.)

Pulsed HIFU facilitated systemic delivery of various agents

Pulsed HIFU enhanced delivery of MR contrast

Among the first studies to demonstrate the feasibility of enhanced delivery using pulsed HIFU without cavitation enhancement was done using MRI as an imaging modality (1.5 T) [13]. A transducer with 10 cm diameter operating at 1.56 MHz frequency and approximately 2 mm diameter spot size and focal length of 7.7 cm (1600× focusing factor; f/# ∼1.3) was used for this study. Three rabbits were treated with a low power, long pulse HIFU exposure at less than 35 W, 10% duty cycle, and 1 Hz repetition rate (100 ms on, 900 ms off) for 5 min. A fourth rabbit was treated with a high power, short pulse HIFU exposure, with maximium 500 W power, 0.1% duty cycle and 1 Hz repetition rate (0.1 ms on, 999.9 ms off), also for 5 min. The transducer position was fixed during the entire treatment time. The powers described in [13] are nominal electric powers, and actual acoustic powers might be somewhat, or in the second instance, substantially less than the figures quoted. Care must be taken when comparing to studies which use TAP. T1 and T2 weighted MR images were obtained prior to treatment, following treatment and after injection of a polymerized liposome particle incorporating an MR (T1) contrast agent (Magnevist, Berlex Laboratories, Wayne, NJ). The nominal size of these particles was 100 nm. For both treatment protocols, bright spots observed on the T2 weighted images one hour after HIFU treatment indicated the presence of edema, confirmed by transmission electron microscopy after biopsy. T1 weighted images following injection of the contrast agent (2.5 hours after treatment) showed increased leakage in the same region. The results for the two different treatment protocols were similar, but the T1 enhancement was apparently greater in the 500 W case. The rabbits were imaged again a week later to show that the edema had disappeared.

Pulsed HIFU enhanced tumor growth inhibition by bortezomib in a murine squamous cell carcinoma model

This study was performed to test the hypothesis that effective tumor growth inhibition can be achieved by combining pulsed HIFU with a dose of the anti-cancer drug, bortezomib [54], which is ineffective when given alone in a murine head & neck squamous cell carcinoma model [55]. Murine squamous cell carcinomas (SCC7) were grown in the right flank of C3H mice. The mice were then randomly assigned into six treatment groups: saline injection control, ‘low dose’ (1.0 mg/kg) bortezomib and ‘high dose’ (1.5 mg/kg) bortezemib, each with and without a preceding pulsed HIFU treatment. Treatments were twice weekly, using the Sonoblate 500 system at 40 W TAP, and pulse parameters of 5% duty cycle, 1 Hz repetition rate (50 ms on, 950 ms off), and 100 pulses per site. The entire tumor volume was treated by rastering. The tumor volume measurement was calculated from the length, width, and height as measured with digital calipers. The starting point for the study was the day the tumor reached a volume of 125 mm³ and the endpoint when it reached 650 mm³. A second set of tumors, without the high drug dose groups, were also grown and treated in the same manner, except that this set was sampled by sacrificing animals on days 1, 3, and 6 (3 animals per group per day). Tumors were then extracted for histology. A statistical comparison between groups was performed on the number of days required to grow to 650 mm³, and a histological apoptotic index based on TUNEL (terminal deoxynucleotidyl transferace-mediated deoxyuridine triphosphate nick end labeling) staining.

The results showed that neither the pulsed HIFU alone, nor the 1 mg/kg dose alone had any impact on tumor growth, however, in combination they resulted in the same statistically significant tumor growth inhibition (p = 0.018, n = 5 and 6) as the 1.5 mg/kg dose alone. The effects were not strictly additive in that there was no further growth inhibition obtained by combining the pulsed HIFU with the high drug dose. The results were corroborated with a significantly greater apoptotic index for the pulsed HIFU plus drug group relative to the others, in animals sacrificed 24 and 72 hours post-treatment.

The study demonstrated that pulsed HIFU and bortezomib have a synergistic effect, but it remains unclear just what the nature of this synergism is. Bortezomib is not a large molecule, having a molecular weight of only 384 Daltons. Therefore, it is less obvious that enhanced delivery is the main effect of the pulsed HIFU. It is also known that bortezomib blocks the molecular pathways that allow tumor cells to respond to stresses such as those that are possibly caused by pulsed HIFU in this case. Further study would be required to determine which of these effects (enhanced delivery of the bortezemib or increased susceptibility to stress) is more important.

Pulsed HIFU enhanced delivery of radiolabeled monoclonal antibody B3 to A431 tumor in nude mice

This study tested the hypothesis that pulsed HIFU could improve the delivery and tumor targeting of a radioactive monoclonal antibody to a mouse tumor xenograft system [56]. The B3 antibody and A431 carcinoma system was chosen because previous work had demonstrated some tumor targeting, but poor antitumor effectiveness, presumably because the systemic dose became toxic before the tumor dose was high enough [57].

Groups of nude mice (n = 5 per time point) were inoculated subcutaneously with 3 × 10⁶ A431 tumor cells in each hind flank. One week after inoculation, one of each tumor pair was treated with pulsed HIFU; while the second tumor served as an untreated control. Two pulsed HIFU power levels were tried using the Sonoblate system, 20 W and 40 W TAP. The pulsing parameters were 5% duty cycle and 1 Hz repetition rate, with 100 pulses per raster site.

Following pulsed HIFU exposures, the mice were injected intravenously with ¹¹¹In-MX-B3. Static imaging was performed on anaesthetized mice using a single photon emission computed tomography system (A-SPECT, Gamma Medica Instruments, Northridge, CA) at time points of 1, 24, 48 and 120 hrs following injection (Figure 1). Immediately following imaging, a number of animals were euthanized, and tissue samples removed for histology. Alternatively, at these same time points, mice were euthanized and dissected, and the decay-corrected radioactivity of the tumors and various organs were measured with a gamma counter (Wallac Inc., PerkinElmer, Inc., Boston, MA). The percentage of injected dose per gram (%ID/g), normalized to a 20 g mouse, could then be calculated for each organ. A plot of %ID/g as a function of time was made for each organ, and the areas under the curve calculated for treated and untreated tumors (Figure 2).

Figure 1. Images of A431 tumor-bearing nude mice receiving intravenous ¹¹¹In-MX-B3. Tumors were exposed to pulsed-HIFU with 2 exposure parameters: TAP of 20 W (A) and TAP of 40 W (B); pulse repetition frequency, 1 Hz; duty cycle, 5% (50 ms on and 950 ms off). One hundred pulses were given at each raster point. Within 10 min, mice were injected intravenously with 11 MBq (32 µg) of ¹¹¹In-MX-B3. Static imaging was acquired for 40,000 counts at 1, 24, and 120 h after injection and immediate anesthetization with ketamine (60 mg/kg)/xylazine (10 mg/kg). Pulsed HIFU–exposed tumors are indicated by arrows while the contralateral, untreated tumors were used as controls. Animals pulsed at 40 W had earlier and higher uptake than those treated at 20 W. The difference between 40 W treated and untreated tumors was not measurable by 120 h. ([56] Reprinted with permission.)

Figure 2. Biodistribution (%ID/g; n = 4–6) of ¹¹¹In-MX-B3 in A431 tumor-bearing nude mice. One tumor was treated with pulsed HIFU at 40 W and the contralateral tumor was used as a control. Each mouse received 74 kBq/1.5 µg of ¹¹¹In-MX-B3 and uptake was measured at 1, 24, 48, and 120 h after injection. Data are shown as%ID/g of tissue and were normalized to a 20-g mouse (mean±6 SD). Insertion is uptake and retention of ¹¹¹In-MX-B3 in A431 tumors. Pulsed HIFU exposure shortened peak tumor uptake time and increased peak tumor uptake value compared with that of untreated control tumors. AUC calculation for 120-h period resulted in a 1.4× higher value for pulsed HIFU-exposed tumors than that for control tumors. * P < 0.05. ([56] Reprinted with permission.)

At 20 W, there was no noticeable difference in tumor uptake between the HIFU and the control tumors when investigated by nuclear imaging at 1, 24 and 120 hr after injection. In contrast, the pulsed HIFU exposures at 40 W resulted in earlier and more intense visualization of tumor than the control without the pulsed HIFU. The remaining studies were carried out using the 40 W TAP only. The highest level (37.55 ± 16.37% ID/g) of ¹¹¹In-MX-B3 in the pulsed HIFU exposed tumors was observed at 24 hr whereas the highest tumor uptake (24.99 ± 2.40% ID/g) in the control tumors without the pulsed HIFU exposure was observed at 48 hr. The uptake in tumors with the pulsed HIFU exposure was more than twice that in the control tumors at 1 (2.57; p < 0.05) and 24 hr (2.13; p < 0.02), but the uptake in the pulsed HIFU exposed tumors decreased almost to the level of the control tumors at 120 hr. The area-under-curve (AUC) analysis for the tumor uptake and retention curves resulted in 3433 ± 886 (%ID × hr/g) and 2521 ± 352 (%ID × hr/g), for the HIFU exposed and control tumors, respectively, over the entire 120 hr period (p < 0.04). The study therefore demonstrated that a single pulsed HIFU treatment could increase the radiation exposure dose of tumors by 36% in this time period, compared to untreated control tumors. Histology did not reveal any tissue damage associated with the pulsed HIFU exposures.

Pulsed HIFU facilitated systemic delivery of naked DNA in a squamous cell carcinoma model

To demonstrate the feasibility of using pulsed HIFU to enhance gene transfection, the systemic delivery of a reporter gene was shown in a tumor model [58]. Squamous cell carcinoma (SCC7) tumors were induced subcutaneously in both flanks of female C3H mice (n = 3) and allowed to grow to an average size of 0.4 cm³. One tumor was exposed to pulsed HIFU using the previously described Sonoblate system, while a second tumor served as a control. Settings for the exposure were as follows: total acoustic power of 20.5 W, total of 120 pulses per site at pulse repitition rate of 0.5 Hz and duty cycle of 4.5% (90 ms on, 1910 ms off). Immediately after ultrasound exposure, a solution containing a cytomegalovirus-green fluorescent protein (GFP) reporter gene construct (5 µg/g) was injected intravenously via the tail vein. The mouse was sacrificed 24 hours later. Each of the six tumors was sliced longitudinally into three approximately equal sections. One slice was used for histopathologic analysis, the second was used for fluorescence microscopy, and the third was used for Western blot analysis. Observations with light microscopy in tumor sections excised at 24 hours post-treatment and stained with Hematoxylin-eosin (H&E) stain indicated that pulsed HIFU did not produce any acute destruction of the tissues. There were also no thermal lesions, hemorrhage or coagulation necrosis. The results of the TUNEL assay did not show a noticeable increase in the number of apoptotic cells in the treated tumors compared with the control tumors. This result indicated that the ultrasound exposures did not cause cell death in the tumors. GFP expression was observed in all three sonicated tumors. Fluorescence was visible in the SCC7 epithelial cells predominantly surrounding blood vessels (Figure 3). Fluorescence was not observed in control (non-exposed) tumors. Quantification of signal intensity measurement at Western blot analysis showed that, at 24 hours post treatment, GFP expression levels were significantly greater in tumors that received exposures (160.2 arbitrary units [au] ±24.5) than in controls (17.4 au ±11.8) (p = 0.004, n = 3) (Figure 4). These results clearly demonstrate better transfection due to pulsed HIFU pre-treatment. This was interpreted as being due to increased transport of the plasmid from the vasculature and through the interstitial spaces.

Figure 3. Representative histologic sections from, (A) control tumor and, (B) tumor treated with pulsed high-intensity focused ultrasound, viewed at fluorescence microscopy. Expression of GFP reporter gene (green) is visible in tumor that underwent ultrasound exposure but not in control. DAPI staining (blue) indicates nuclei of tumor cells. Bar =50 µm. (Original magnification ×100.) ([58] Reprinted with permission.)

Figure 4. (A) Western blot analysis shows expression of GFP gene (top row) against that of α-actin (bottom row) in tissue samples from control tumors (a, c, e) and treated tumors (b, d, f). (B) Graph of results of densitometric analysis shows significant increase in GFP expression in tumors treated with pulsed high-intensity focused ultrasound (160.2 arbitrary units [au] ±24.5) compared with that in control tumors (17.4 au ± 11. 8) (paired Student t-test, P = 0.004). Values were normalized to those of the housekeeping gene, α-actin. ([58] Reprinted with permission.)

Pulsed HIFU enhanced viral gene delivery in a mouse tumor model

Viral gene delivery was also demonstrated in a mouse tumor model [59]. Human glioblastoma multiforme LN229 tumors were implanted on both hind flanks of nude mice (Balb/C nu/nu). One side was treated with the Sonoblate 500 pulsed HIFU using an exposure regimen of 5% duty cycle, 1 Hz repetition rate (50 ms on, 950 ms off) at 100 pulses per site. Following treatment, adenovirus AD-cmv-Luc was injected via the tail vein. In this study, six different total acoustic power (TAP) levels were evaluated, 10 W, 15 W, 20 W, 25 W, 30 W and 40 W, with 5 mice in each group. Bioluminescent imaging was conducted at 48 hours post-treatment on the nude mice, immediately followed by euthanasia of the animals and extraction and processing of tumors, livers, and lungs. In vitro analysis included Western blotting and immunohistochemistry. The bioluminescent intensity was compared for identically sized regions of interest (5 mm diameter circle) of the treated and control tumors. Only the 20 W treatment showed an enhancement. This was confirmed by the in vitro assay (treated: 121+/−73 cps/g versus control: 33+/−9 cps/g of luciferase activity). At all other TAP levels the luciferase expression levels were lower than the control group. Only the 10 W level was not statistically different from the controls. The follow up histology using H & E and TUNEL staining showed that, for exposures above 20 W, the tumors became damaged and necrotic, while at 20 W and below, no clear damage was seen. This appears to contradict the previously mentioned study [58] that was performed at 40 W TAP, however, it is possible that this tumor type is more sensitive than the SCC7 tumors. Transmission electron microscopy performed on 20 W and control samples appeared to indicate limited damage to the endothelium, as well as changes in endothelial junctions in the treated samples.

Pulsed HIFU delivery of Doxil and Thermodox

Doxorubicin is a potent anti-cancer drug that researchers have been seeking to better target to tumors [60], [61]. Doxil® (Ortho Biotech, Inc.) and Thermodox™ (Celsion Corp.) arose out of progressively more sophisticated attempts to target the drug by encapsulating it in liposomes of approximately 100 nm diameter. The former targets passively based on size alone, while the latter actively releases its payload when heated. Two studies were conducted to determine whether pulsed HIFU might enhance the delivery and therapeutic effectiveness of Doxil, with somewhat contradictory results. In the initial study [62], SCC7 tumors were used, grown in the flanks of C3H mice. A custom 1.5 MHz transducer (as in [13]) was employed for the ultrasound exposures. The exposures in this case were 5 min long, targeting a single point at the center of the tumor, at a nominal (electrical) power of 35 W, a repetition rate of 0.5 Hz and a series of duty cycles from 2.5% (50 ms pulse length) to 6% (120 ms). Unlike most other studies detailed herein, the therapeutic agent in this case was injected prior to HIFU treatment. Mice were sacrificed 2 hr after treatment, following which the tumors were extracted and assayed for doxorubicin content. The greatest drug concentration was demonstrated at 4% duty cycle (80 ms pulse length), so, only at this setting were compared control, Doxil only, and Doxil-plus-pulsed HIFU treatment groups. In that case, a 124% increase in doxorubicin concentration was demonstrated in the tumors that had received HIFU compared to those that were given the drug by itself (p < 0.001, n > 5) (Figure 5).

Figure 5. Graph shows mean doxorubicin content in tumors in mice in group 1 (pulsed HIFU and Doxil), group 2 (Doxil without pulsed HIFU), and group 3 (controls, no Doxil or pulsed HIFU). Analysis of group 3 tumors showed no peaks at the major emission frequencies for doxorubicin and its metabolites. Group 1 tumors showed a 124% (9.4 µg · g⁻¹/4.2 µg · g⁻¹) increase in free doxorubicin over that in group 2 tumors (P < 0.001). ([62] Reprinted with permission.)

The second study [63] was similar to the first but with key differences. A murine mammary adenocarcinoma cell line (JC) tumor was grown in both flanks of each female syngenic Balb/c mouse. SCC7 tumors, previously described, were also tested. In this case, the Sonoblate pulsed HIFU system was used, with a TAP of 20.5 W, with a 9% duty cycle and 1 Hz repetition rate (90 ms on, 910 ms off), with 120 pulses at each raster point. Following pulsed HIFU treatment (unlike the previous study), the Doxil was administered via tail vein injection. Tumor volume was monitored by measuring with calipers for groups of mice receiving 0 1, 2, 6, and 10 mg/kg of Doxil, with and without adjuvant HIFU. Also, tumors were assayed for doxorubicin content as in the earlier experiment but at a greater number of timepoints, starting at 5 min following injection and extending out to 180 min. No significant differences could be found, either in tumor growth or in the Doxil assay, between tumors treated with Doxil alone and those pretreated with pulsed HIFU. This was found to be the case for both the JC and the SCC7 tumor models, despite a difference in the mean uptake of Doxil between them. It was hypothesized that due to the long circulation time of Doxil, the time points for performing the Doxil assay were not long enough to correlate with the therapeutic response.

Studies of fluorescent agent delivery, also reported in these two papers, were in qualitative agreement. The first paper reports on an adjuvant study showing how the delivery of a 500 kDa dextran–fluorescein isothyocyanate (FITC-dextran) is increased with increasing number of pulses (from 30 to 150) of pulsed HIFU (Figure 6), while the second discussed a similar effect in muscle, as well the enhanced delivery of 100 nm fluorescent microspheres (Molecular Probes, Eugene, OR) to the JC tumors.

Figure 6. Representative confocal micrographs of the tumors in the mechanistic study. (A, C) Images of control tumors of mice treated with 30 and 150 pulses per site, respectively. (B, D) Images of tumors treated with 30 and 150 pulses per site, respectively. All other high intensity focused ultrasound parameters were constant: frequency, 1.0 MHz; pulse repetition frequency, 0.5 Hz; spatial average–time average intensity, 1376 W/cm²; and time, 2 min per site. Increased extravasation is evident in tumors that were exposed to high-intensity focused ultrasound. Relative increase in extravasation is also evident in D, where 150 pulses per site were given, as opposed to B, which received 30 pulses per site. Permeability appears to be isolated to capillary beds and not related to larger vessels. Large arrows = blood vessels, small arrows = postcapillary venules, bar = 80 µm. ([62] Reprinted with permission.)

In a follow-up study, pulsed HIFU was used as a thermal trigger for deploying doxorubicin from Thermodox [33], which is a low temperature heat sensitive liposome formulation designed to break down at temperatures around 42°C. In this case, the agent was systemically injected prior to HIFU treatment, other than that, the treatment parameters and assay methods were very similar to those for the previous Doxil study. Both the tumor growth and the doxorubicin retention in the treated tumors showed a significant improvement (p < 0.05, n = 5) over the untreated contralateral control [33]. The paper shows that the time taken for the treated tumor to grow to a volume of 500 mm³ is almost double that of the control (11.8 versus 5.7 days), while the amount of doxorubicin assayed in treated tumors extracted immediately following sonication is more than double that of the control (precise number not given).

Pulsed HIFU facilitated local delivery of agents

One of the unique claims of pulsed HIFU delivery has been that it appears to alter properties not only of the vasculature, but also of the bulk tissue, as represented in Figure 7. This claim has been based mainly on observations [64] that agents injected directly into treated tissue appear to spread over a much wider region than in similar untreated tissue, where injected agents tend to remain in the needle track. While this claim may be based on an overly compartmentalized and simplistic understanding of tissue transport, it does remain a tantalizing possibility, in that no other physical targeting method, including microbubble enhanced HIFU, can make such a claim. Also, for tumors, the major barrier to transport for molecules in the 40–100 nm range is not vascular impermeability, but rather poor convection due to high interstitial pressures (∼8–30 mmHg) [11], [12] and low pressure gradients. The study outlined below demonstrates that this aspect of pulsed HIFU delivery warrants further investigation. It is particularly relevant in light of the fact that so many gene and cell therapy trials use local rather than systemic injections to deliver their agent into tumors, bypassing the vasculature entirely in an effort to avoid systemic toxicity [65]. If pulsed HIFU were proven to assist in dissemination of the genetic vectors, it could help make gene therapy a more practical reality.

Figure 7. Schematic representation of effect of high-intensity focused ultrasound exposures on tumors in conjunction with intravenous injection of liposome-encapsulated doxorubicin (Doxil) as a model. (A) Region of tumor with intact endothelium (E) surrounding vessel (v) and densely packed tumor cells (T). (B) Same region as in A after high-intensity focused ultrasound exposure. It is hypothesized that focused ultrasound energy disrupts endothelial barriers and changes interstitial properties (alters vascular permeability and hydrostatic pressures in the parenchyma), enhancing extravasation and distribution of pharmaceutical agents. ([62] Reprinted with permission.)

Pulsed HIFU enhanced delivery of TNF alpha

Tumor necrosis factor-α (TNF-α) is a cytokine that induces hemorrhagic tumor necrosis and tumor regression. The use of the TNF-α gene for tumor therapy has been studied but with limited success. This is thought to be due to insufficient distribution of the plasmid and dose limiting toxicity at concentrations needed for tumor growth inhibition and regression [66], [67]. The objective of this study was to determine whether pulsed HIFU could enhance the local delivery of TNF-α plasmid and consequently improve its therapeutic effects [68]. SCC7 cells were grown in culture and then injected subcutaneously into shaved bilateral flanks of syngeneic female C3H mice. Two weeks following tumor cell inoculation, the mice were randomly assigned into 6 groups of 5 mice each: control with no injection or HIFU, saline injection only, TNF-α plasmid only, pulsed HIFU only, HIFU with saline and HIFU with TNF-α plasmid. Pulsed HIFU treated tumors were exposed immediately prior to injection using the Sonoblate 500 system. The exposure parameters were 40 W TAP, 1 Hz pulse repetition frequency, and 5% duty cycle. One hundred pulses were given at each raster point. Tumor volume was measured as in previously mentioned studies using calipers. Pulsed HIFU exposures alone, intratumoral injections of saline alone, or the combination of the two did not change the growth rates of tumors relative to untreated controls. Based on the average volume of the tumors 5 days after the treatments, a significant (p < 0.05, n = 5) reduction in tumor growth was observed for tumors receiving direct injection of TNF-α plasmid (40% less volume than control). A further significant reduction in tumor growth was found when the TNF-α injections were preceded by pulsed HIFU exposures (74% less volume than TNF-α plasmid alone).

H & E staining and direct injections of a surrogate agent (200 nm fluorescent microspheres) were used to compare patterns of necrosis and distribution with and without pulsed HIFU exposures. In the pulsed HIFU treated tumors, the spread of fluorescent particles from the needle track was broader and deeper, whereas in the untreated tumors, the particles stayed predominantly along the needle track and on the surface of the tumor (qualitative observation). This result was consistent with the relative sizes of the necrotic regions observed in the pulsed HIFU treated and untreated control tumors, both of which were injected with the TNF-α plasmid. The results would appear to indicate that the improved therapeutic response observed when combining pulsed HIFU with locally administered TNF-α plasmid may be due to the improved diffusion of the plasmid within the solid tumor.

Pulsed HIFU thrombolysis with rtPA

Many reports over the years have suggested that ultrasound might be used for thrombolysis, either alone or in conjunction with cavitation inducing UCAs [69–71]. This study was designed to test the hypothesis that pulsed HIFU might work to improve the effectiveness of the ‘clot busting’ drug recombinant tissue plasminogen activator (rtPA) by improving drug penetration into the clot [16]. An earlier in vitro study [72] had shown that the penetration and breakup of in vitro clots by rtPA was indeed improved following pulsed HIFU treatment. In the follow up in vivo study, a clot model was induced in the marginal ear vein of rabbits (n = 4), where it could be easily accessed with the HIFU. A special platform was designed so that the rabbits could be reclined above the water bath and targeted through a thin membrane with degassed water used as couplant. The pulsed HIFU treatment was similar to that previously described, with 40 W TAP, 5% duty cycle, 1 Hz repetition rate, and 100 pulses per site along the clotted vein. A high frequency animal ultrasound scanner (Vevo 660, Visualsonics, Toronto, Canada) was used to assess the clot integrity at regular intervals following treatment. This device, operating at 55 MHz, could easily distinguish between the static clot tissue and flowing blood in the vein by picking up the backscatter from the moving red blood cells. B-scan images were taken at 6 points along the 3 cm length of the clot, and scored as either closed or open. While neither HIFU nor rtPA alone was capable of opening the clot in a 5-hr period, the synergistic effects of pulsed HIFU and rtPA resulted in the dissolution of the entire clot for every rabbit in that treatment group. Based on the in vitro studies, the authors concluded that the mechanical action of the pulsed HIFU treatment allowed the rtPA to penetrate beyond the surface of the clot and thereby work much more efficiently. A very recent study that similarly coupled ultrasound and rtPA at 1 MHz and 120 kHz came to similar conclusions but with a somewhat different notion of the mechanisms [73]. In that paper, the greatest clot loss was concluded to be coincident with stable cavitation, although the presence of the cavitation was inferred indirectly from the acoustic amplitude rather than being measured.

Studies of pulsed HIFU mechanism

A number of studies have been performed in an attempt to characterize the physical aspects of a pulsed HIFU treatment and thereby elucidate a mechanism. As previously mentioned, the two strongest ultrasound tissue interactions generally recognized by the ultrasound community are heat and cavitation, with radiation forces and other possibilities only being considered if these two are demonstrably absent. In the pulsed HIFU work, a lack of histological indications of thermal or hemorrhagic damage was taken as early evidence that neither heat nor cavitation is an important factor. Indeed, the pulsed HIFU exposure parameters were designed precisely to avoid these two effects, in order to prevent any permanent tissue damage. Since then, a series of studies [63–76] have dealt with the thermal issue. Two of these used embedded thermal probes to measure the temperature rise in mouse tumor [63] (fluorometric probe) and leg muscle [75], [76] (hypodermic thermocouple) during typical pulsed HIFU exposures with parameters of 40 W power, 5% duty cycle, 1 Hz repetition rate, and 100 total pulses per site. In [75] and [76], the peak temperature was calculated based on extrapolating the temperature rise measured at multiple points via a fit of the bio-heat equation. This was done to avoid direct probe heating. The other study used MR thermometry (spatial resolution: 2 mm, temporal resolution 3.9 s) to detect the temperature rise during pulsed HIFU treatment of rabbit thigh muscle, for a series of treatment powers between 10 and 60 W TAP, at 5 and 10% duty cycle, 1 Hz repetition rate, and 120 total pulses per site [74]. The peak temperatures measured during these studies ranged from 2.8°C (at 40 W) [74] in the rabbit thigh, to 5.2°C [63], [75] in the mouse tumor and muscle, above ambient. Tissue remained at this temperature for the duration of the exposure (∼2 min) before returning to the baseline as the heat rapidly diffused out of the focal zone in all directions. Presumably, since these exposure parameters are the same as those in most of the previously mentioned studies, these experimental observations reflect what happened in those studies as well. For comparison, hyperthermia treatments for drug delivery done at these temperatures can last 20–60 min. The most comprehensive study [75], [76] went a step further in replicating the thermal dose (≥42°C for 2 min, plus cool down) using an infrared heat source, and comparing directly its effectiveness in assisting the systemic delivery 200 nm polystyrene fluorescent microspheres with that of the pulsed HIFU treatment period. A histological assay of the particle density was used to quantify the delivery, and the comparison was to the contralateral untreated control limb of the same animal. Only the pulsed HIFU treatment showed a significantly improved particle distribution. Also, it was found that dropping the bath temperature from the usual 37°C (peak treatment temperature ∼42°C) down to 34°C (peak treatment temperature ∼39°C) improved the pulsed HIFU effect by almost 3-fold. Finally, it has been demonstrated that the effects of the pulsed HIFU treatment last at least 24 hours [76]. These three facts, the inability of a similar hyperthermia treatment to duplicate the effect, the increase in delivery despite a reduction in peak temperature, and the long recovery time, are difficult to reconcile with a traditional hyperthermia mechanism. The same study [75], [76] tried, but was unable to unambiguously eliminate or confirm cavitation as a mechanism. Part of the reason for this lies in the animal model. Because the focal zone (7 mm long) is almost the same dimension as a mouse thigh (4–7 mm), it is impossible to avoid serious interactions between the ultrasound and these physical features. The result is that ultrasonic backscatter spectroscopy, the most sensitive and commonly used technique for monitoring cavitation in opaque materials [45], [77], [78], cannot distinguish between the cavitation occurring at the skin or bone surface and that in the muscle itself. While the histology has consistently shown little of the damage, such as hemorrhage and cell lysis, that is most often associated with destructive inertial cavitation [79], more subtle effects, such as edema, might be related to cavitation that is restricted to and impacts only on the vasculature. Even had no cavitation signatures been detected in the spectra, it is still possible to argue that the levels are simply below the threshold for detection. It is certainly well known that pulsed HIFU at sufficient intensities in the focal region can produce cavitation in tissues, even in the absence of UCA microbubbles [79]. Nevertheless, there has been much speculation in this literature about the possibility of alternative mechanisms, principally due to the action of shear stresses from radiation forces [63], [80].

Conclusion

Ultrasound therapies for enhancing the delivery of macromolecular agents remain tantalizing, although poorly understood, medical tools. Considerable research has gone into demonstrating that this form of energy may be useful in breaking down transport barriers at the skin, the vascular, and the cell membrane levels when cavitation is initiated through the use of UCA microbubbles. Pulsed HIFU at slightly higher energies has demonstrated similar promise without the use of UCAs. In that case, however, the mechanism is even more controversial. Hyperthermia alone does not appear able to duplicate the affect, leaving little to support the hypothesis of a purely thermal mechanism. Conversely, it has not been possible to preclude a cavitation based mechanism based on the studies that have been carried out to date. Nevertheless, at least some data, particularly histological data, has been interpreted as supportive of a third possibility, that shear stresses due to radiation force displacements might be sufficient to cause structural changes (‘gaps’) to open up in the tissue. The strains produced in this way are not large, however, so it is unclear what tissue structures they might possibly be targeting. Another possibility, which deserves more attention, is that of a synergistic coupling of the mechanical or thermal stressors to biological pathways capable of producing the observed effects. At least one study explicitly considered this as a possible explanation for the increased efficacy of a small molecule therapy (bortezemib) when coupled with pulsed HIFU [54]. In that case, as detailed above, it is thought that the bortezemib and the pulsed HIFU target the same cellular pathways. It is also possible that the improved delivery itself might be primarily due to biology. For example, edema has been observed to accompany the improved delivery in at least two studies [13], [75]. While this may be an effect of pulsed HIFU on the vasculature, it is also possible that the local stresses due to HIFU treatment might lead to a release of histamines, resulting in a systemic inflammatory response, as the body tries to contain the ‘damage’. Biological stress responses have been implicated in other physical drug delivery techniques, including hyperthermia [1] and even microbubble enhanced delivery [81]. Exploring all these possible interactions is important not only for scientific understanding of the process and its safety margins, but also because of treatment-drug interactions that may arise.

Declaration of interest: The authors report no conflicts of interest. The authors alone are responsible for the content and writing of the paper.