Multifocused ultrasound therapy for controlled microvascular permeabilization and improved drug delivery

Abstract

Focused ultrasound (FUS) exposure of microbubble (MB) contrast agents can transiently increase microvascular permeability allowing anticancer drugs to extravasate into a targeted tumor tissue. Either fixed or mechanically steered in space, most studies to date have used a single element focused transducer to deliver the ultrasound (US) energy. The goal of this study was to investigate various multi-FUS strategies implemented on a programmable US scanner (Vantage 256, Verasonics Inc) equipped with a linear array for image guidance and a 128-element therapy transducer (HIFUPlex-06, Sonic Concepts). The multi-FUS strategies include multi-FUS with sequential excitation (multi-FUS-SE) and multi-FUS with temporal sequential excitation (multi-FUS-TSE) and were compared to single-FUS and sham treatment. This study was performed using athymic mice implanted with breast cancer cells (N = 20). FUS therapy experiments were performed for 10 min after a solution containing MBs (Definity, Lantheus Medical Imaging Inc) and near infrared (NIR, surrogate drug) dye were injected via the tail vein. Fluorescent signal was monitored using an in vivo optical imaging system (Pearl Trilogy, LI-COR) to quantify intratumoral dye accumulation at baseline and again at 0.1, 24, and 48 h after receiving US therapy. Animals were then euthanized for ex vivo dye extraction analysis. At 48 h, fluorescent tracer accumulation within the tumor space for the multi-FUS-TSE therapy group animals was found to be 67.3%, 50.3%, and 36.2% higher when compared to sham, single-FUS, and multi-FUS-SE therapy group measures, respectively. Also, dye extraction and fluorescence measurements from excised tumor tissue found increases of 243.2%, 163.1%, and 68.1% for the multi-FUS-TSE group compared to sham, single-FUS, and multi-FUS-SE therapy group measures, respectively. In summary, experimental results revealed that for a multi-FUS sequence, increased microvascular permeability was considerably influenced by both the spatial and temporal aspects of the applied US therapy.

I. Introduction

Cancer is a public health problem and is the second most common cause of death in the United States [1]. Neoadjuvant therapy has a major role in the treatment of cancer and includes chemical, radiation, and hormone therapies. Among these, chemotherapy is the first line of treatment for most cancer types. However, it has been suggested that only about 1% of the injected drug dose reaches the tumor site [2], [3]. Local delivery and penetration of the chemotherapeutic drug is limited by various factors associated with the tumor microenvironment. One of these factors is the presence of an abnormal microvascular network. Some of these tumor vessels are chaotic and leaky, which increases the interstitial fluid pressure resulting in poor drug penetration of tumor tissue [4]. To improve drug penetration, increasing the overall drug dose is generally not a viable option due to concerns of systemic toxicity. To avoid this adverse effect, different types of targeted drugs and carriers have been developed in the past few decades. For example, microbubble (MB) or liposomes can be injected systemically to safely transport loaded drugs or genes to the tumor space whereby the cargo is noninvasively released by ultrasound (US) exposure.

It has been previously reported that focused US (FUS) can safely and reversibly increase microvascular permeability and improve delivery of an intravenously injected drug to targeted cancer tissue [5], [6]. Exposure of MB contrast agents to low-intensity US is used to help control this process and to reduce the risk of tissue damage [7]. US at lower acoustic pressures (order of a few hundred kilopascals, kPa) can cause circulating MBs to volumetrically expand and contract in response to the rarefaction and compression phases of the pulsed US waves, respectively. Termed stable cavitation, this process transiently disrupts microvascular structures and helps tumor uptake of a circulating drug. Typical peak negative pressure values that produce stable MB cavitation are in the range of 0.3 to 2.3 MPa for a 2.0 MHz US pulse [8]. It has been shown that lower US pressures (< 1.5 MPa) can successfully induce microvascular permeabilization due to stable cavitation, whereas higher US pressure levels result in inertial cavitation and MB destruction.

Several studies have shown that US-mediated systemic delivery of drug or adenoviral vectors improves uptake at the targeted tissue space in animal models of cancer [9]–[15]. Only recently, Kotopoulis et al. [16] demonstrated the ability and efficacy of FUS-mediated gemcitabine drug delivery in a clinical setting while using commercially available MBs. Results suggested that patients with inoperable pancreatic cancer had prolonged survival due to reduced tumor size and growth. This pioneering study represents the first time that US-mediated drug delivery has been used in human patients to improve anticancer treatment. In addition, numerous studies have shown you can locally disrupt the blood-brain barrier (BBB) to transiently facilitate targeted drug delivery [17]–[21].

Most FUS-mediated drug delivery studies to date have relied on use of a single focal zone from a fixed US transducer to improve microvascular permeabilization. However, a single US focus only allows for MB interaction and US therapy within the beamwidth at the focus. Consequently, this strategy fails to effectively treat any tissue outside the US focus. One can move the transducer manually or mechanically over the tumor to help treat the entire cancer volume [11], [22]–[24]. However, the mechanical movement of the transducer can lead to poor coupling between the transducer/skin interface and require more time to move to the next focal zone, which increases overall treatment time [25], [26]. It is also difficult to maintain a uniform US energy distribution during manual movement of the transducer. Recently, it was shown that a multi-element array transducer could be used for microvascular permeabilization in volume space by beam steering during US exposure [27]–[29]. As described by Huang et al. [27], an US beam was steered over a 3 × 3 rectangular grid that included nine focal locations spaced 3 mm apart. Each location received pulsed US for 300 ms (2 msec on, 28 msec off) for a total exposure time of 50 sec. This led to successful BBB openings of approximately 1 cm³ in volume. A similar US pulsing sequence was later used to transiently improve microvascular permeabilization in animal models of Alzheimer’s disease [28] and clinically in the primary motor cortex of patients with amyotrophic lateral sclerosis [29]. These studies used an US pressure ramp test to determine the inertial cavitation threshold and then used the same parameters during US therapy. However, there is no detailed study on the sequencing of US focusing and the therapeutic impact. There have also been no studies to date that have reported on the use of a multi-FUS strategy applied over an entire tumor region to improve US image-guided drug delivery.

In this paper, we introduce a multi-FUS system and method to further improve the US therapeutic process. Specifically, we introduce different strategies to perform multi-FUS therapy based on the excitation sequence of user-defined focal zones. Multi-FUS therapy approaches are compared to single-FUS and sham therapy. Preliminary in vivo experiments were performed using an animal model of breast cancer. A small molecule near infrared (NIR) dye was used as a surrogate drug and allows for optical localization and longitudinal tracking in live animals. Both in vivo optical imaging and ex vivo extraction methods were used to quantify dye accumulation in the targeted tumor tissue.

II. Materials and Methods

The US image-guided multi-FUS therapy system utilizes several focal zones spread over a defined tumor area. Fig. 1 illustrates the US image-guided therapy setup for both single-FUS and multi-FUS approaches. Single-FUS therapy uses a fixed focus, whereas multi-FUS therapy uses beamsteering and many focal zones spread over the entire tumor space. In this research, we implemented and investigated the impact of spatially distributed focal zones within a single plane. Multi-FUS therapy is implemented in two ways:

1. Multi-FUS with sequential excitation (multi-FUS-SE): Different focal zones within the tumor region are targeted one at a time in a sequenced fashion, where this sequential targeting is repeated M times (with the same order) for a treatment time at each focal location of N sec.

2. Multi-FUS with temporal sequential excitation (multi-FUS-TSE): Each focal zone within the tumor region is targeted one at a time and in a repeated fashion (defined as repetition per focus). After targeting of one focal point, a new focal point within the tumor region is repeatedly targeted for a predetermined treatment time.

A diagram of these two different FUS treatment sequences is shown in Fig. 2.

Fig. 1.

(A) Illustration of the ultrasound (US) image-guided therapy setup for both single and multi-focused US (FUS) strategies. (B) Schematic represents the excitation sequence for multi-FUS-SE and multi-FUS-TSE approaches. The red colored focal zone represents the active excitation zone.

Fig. 2.

Treatment sequence (pulsing) diagrams for both (A) multi-FUS-SE and (B) multi-FUS-TSE, where M and N denote the number of focal zones and length of treatment time, respectively.

A. Ultrasound Therapy System

An US image-guided FUS therapy system was used for this study [6]. This system consisted of a programmable US research scanner (Vantage 256, Verasonics Inc, Kirkland, WA) equipped with a dual transducer configuration (HIFUPlex-06, Sonic Concepts, Bothell, WA). This co-registered device included a focused transducer for therapy applications and a linear array transducer for interleaved US imaging and guidance. Both the therapeutic and imaging transducers were 128-element arrays with center frequencies of 2.0 and 3.5 MHz, respectively. The latter array elements are arranged in a concentric configuration that allows for beam steering in volume space. A graphical user interface (GUI) on the Vantage system allows for visualization of the target tissue and therapeutic control with protocol design and preview software features. The treatment area was chosen from the B-mode US images and was dependent on size of the tumor. Therapeutic US exposure involved a 1000 msec repetition rate (400 msec on, 600 ms off) applied for a duration of 10 min (600 pulsed US exposures) and using a mechanical index (MI) of 0.5 [9]. A single focal spot was used for the single-FUS sequence, whereas a number of focal locations were used during application of multi-FUS and dependent on the tumor size and focus overlap. A fixed 50% overlap between the focal spots was used both axially and laterally. Repetition per focus was the repeat of one full cycle of US at a particular focal position before steering to the next focal location. A single repetition per focus was used in single-FUS and multi-FUS-SE methods, whereas a repetition per focus was calculated as the ratio of the total number of repeat US exposures to number of focal zones used for the multi-FUS-TSE therapy. For example, if the number of focal zones was set to 40, the repetition per focus was set to 15.

B. Temperature Monitoring

To study the temperature rise due to the proposed US pulse sequences, we prepared a tissue-mimicking phantom material [30] and used a digital thermometer (CL3512A, Omega Engineering Inc, Norwalk, CT) to monitor any temperature change during US exposure. The thermometer was inserted into the phantom and placed at the location of the US focus.

C. Acoustic Measurements

A calibrated hydrophone system (AIMS III, Onda Corp, Sunnyvale, CA) was used to measure the acoustic output and beampattern of the therapeutic US transducer. This system consisted of a degassed tank of water and stepper motors for positioning and control of a calibrated hydrophone (Model HGL-0400, Onda Corp, Sunnyvale, CA) in 3-dimensional (3D) space. This hydrophone was connected in series with a preamplifier and digital oscilloscope. As the hydrophone was mechanically scanned within the US transducer field-of-view, system software recorded voltage signals, then converted and mapped the US property of interest, e.g. peak negative pressure.

D. Animal Preparation

Animal experiments were reviewed and approved by the Institutional Animal Care and Use Committee (IACUC) at the University of Texas at Dallas. Six-week-old female athymic nude mice (N = 20, Charles River Laboratories, Wilmington, MA) were implanted subcutaneously with one million breast cancer cells (MDA-MB-231/Luc, Cell Biolabs, San Diego, CA) in the inguinal region of the mammary fatty pad. The implanted tumors were allowed to grow for at least four weeks. Animals in all groups developed comparably sized tumors ranging from 500 to 610 mm³. Mice were randomly sorted into four groups (N = 5 each): sham therapy (without US), single-FUS therapy, multi-FUS-SE therapy, and multi-FUS-TSE therapy. During FUS therapy, animals were placed on a heating pad to maintain core temperature (~37 °C) and were anesthetized with 1 to 2% isoflurane in oxygen. Prior to the application of any US image-guided FUS therapy, a catheter was securely placed in the tail vein for all intravenous injections. A saline solution containing MB contrast agent (2.3 × 10⁷ MBs; Definity, Lantheus Medical Imaging, Inc, N Billerica, MA) and NIR dye (50 μg; IR-780, Sigma-Aldrich, St. Louis, MO) was then slowly injected via the placed catheter (approximately 5 s). Immediately thereafter, a single session of each FUS therapy was then applied to the tumor tissue of the respective animal groups. The water-backed polystyrene-coated acoustic aperture with US transmission gel (Aquasonic 100, Parker Laboratories, Fairfield, NJ) was used to reduce the air gap between the tumor and US transducer. The same procedure was used to apply sham therapy albeit without any active US exposure.

E. In vivo Optical Imaging

Optical imaging allows longitudinal measurement of local fluorescent tracer uptake and any US therapy-induced effects in live animals [12]. Images were acquired at baseline and again at 0.1, 24, and 48 h following application of FUS therapy. Imaging was performed using a small animal optical imaging system (Pearl Trilogy, LI-COR Bioscience, Lincoln, NE). The fluorescence imager was operated using an 800-nm channel excited at a wavelength of 785 nm with an emission filter of 820 nm. In addition, white light digital images were acquired. The fluorescent signal intensity was measured within a user-defined region-of-interest (ROI) using system software (Image Studio Software, LI-COR Biosciences). The tumor was located using guidance of the white light image and then the ROI was manually drawn to encompass the entire tumor region. The mean fluorescent signal was then calculated from the co-registered ROI. All measurements were first normalized to the background signal and then by ROI pixel count to quantify mean intratumoral fluorescent signal activity, which was used as a surrogate measure of drug delivery and tumor accumulation.

F. Tumor Tissue Dye Extraction

Following the optical imaging session at 48 h, animals were humanely euthanized and tumor tissue was excised and used for a dye extraction protocol [11], [12]. Tumors were cut into smaller pieces and rinsed with saline. These pieces were mixed with 1 mL of radioimmunoprecipitation assay buffer and transferred to 2 mL tubes containing ceramic beads. The buffer had 50 mM Tris-base (pH 7.4), 150 mM sodium chloride (NaCl), 1% Triton X-100, 0.5% sodium deoxycholate, and 0.1% sodium dodecyl sulfate (SDS) as previously described [11]. The tubes then underwent high force homogenization (Bead Mill 4 Homogenizer, Fisher Scientific, Waltham, MA) for IR-780 dye extraction. After centrifugation at 2000 g for 10 min (repeated twice), the supernatants were transferred to a 96-well black plate (200 μL per well). Controls of known IR-780 concentration were processed along with the supernatants of the tissue samples. Each tissue sample was measured in triplicate. The fluorescent signal from each well was quantified by a microplate reader (Synergy H4, BioTek, Winooski, VT) with optical excitation and emission set at 780 nm and 820 nm, respectively. NIR dye accumulation in the tumor tissue was calculated and represented as the percentage of dye retained in the tumor relative to the total dye injected into the animal.

G. Statistical Analysis

All experimental data was summarized as the mean ± standard error (SE). A repeated-measures one-way analysis of variance (ANOVA) and two-way ANOVA tests were used to assess longitudinal trends in the ex vivo dye extraction data and in vivo optical imaging group data, respectively. A p-value of less than 0.05 was considered statistically significant. All statistical analyses were completed using Prism 8.3 (GraphPad Software, Inc, San Diego, CA).

III. Results

The measured 3D pressure field from the FUS therapy transducer is shown in Fig. 3. These orthogonal images highlight the US beampattern and peak negative pressure measurements. The focal size was calculated by estimating the full-width at half maximum (FWHM) and found to be 8.0 × 0.8 × 1.2 mm (depth × lateral × azimuth). For a fixed US output, the peak negative pressure was measured to be 0.7 MPa (MI = 0.5). Temperature monitoring in a phantom material found a 0.3 °C increase at the US focus when excited with a single-FUS sequence and an increase of 0.2 °C when excited with either the multi-FUS-SE or multi-FUS-TSE sequences.

Fig. 3.

Measured beampattern of the FUS therapy transducer.

Longitudinal sequences of in vivo optical images detailing intratumoral fluorescent tracer uptake for the single-FUS, multi-FUS-SE, multi-FUS-TSE, and sham therapy at 0 (baseline), 0.1, 24, and 48 h are depicted in Fig. 4. A review of these image sequences suggests that NIR dye accumulation in the treated tumor tissue is more pronounced after application of multi-focus-TSE therapy compared to the other US therapy (exposure) strategies. Quantification of the fluorescent signal is plotted in Fig. 5 and temporally compare NIR dye uptake in tumor tissue for the single and multi-FUS therapies. Longitudinal measurements revealed a 46.0% increase in fluorescent tracer accumulation for the multi-FUS-TSE therapy over sham therapy measurements at 0.1 h, followed by increases of 64.3 and 67.3% at 24 and 48 h, respectively. The multi-FUS-TSE datatrended towards statistical significance when compared with sham therapy (p = 0.095). Optical measurements from the multi-FUS-TSE therapy group animals were found to be 50.3 and 36.2% higher when compared to single-FUS (p = 0.22) and multi-FUS-SE (p = 0.28) therapy group findings, respectively. At 48 h, a 11.3% increase in the fluorescent signal was found when using the single-FUS therapy (compared to control measures) and increases of 22.9% and 67.3% when employing the multi-FUS-SE and multi-FUS-TSE approaches, respectively. Of note, there was only a slight difference between the single-FUS and multi-FUS-SE therapy approaches and optical images detailing intratumoral NIR dye accumulation was not significantly improved over sham therapy (p > 0.05).

Fig. 4.

Representative in vivo optical images of breast-tumor bearing animals of sham (control, row one), single-FUS (row two), multi-FUS-SE (row three), and multi-FUS-TSE (row four) after the injection of a surrogate drug (near infrared, NIR, dye). Fluorescence images were acquired at baseline immediately before application of US therapy (baseline) and again at 0.1 h, 24 h, and 48 h.

Fig. 5.

Summary of in vivo optical imaging results from breast tumor-bearing animals exposed to single-FUS, multi-FUS-SE, and multi-FUS-TSE therapy compared to sham US therapy at 0.1, 24, and 48 h. Note the improved fluorescent dye uptake using the multi-FUS-TSE therapy protocol.

Excised tumor tissue samples were processed to extract the NIR dye, which provides an ex vivo measure of accumulation and a reference standard, Fig. 6. Fluorescent signals from the dye extraction protocol revealed a significantly higher uptake when using multi-FUS-TSE therapy (p < 0.05). Specifically, the multi-FUS-TSE therapy approach exhibited a 3-fold increase in the measured fluorescent signal compared to tumors that received sham US therapy. These results indicate that NIR dye extracted from excised tumor samples is similar to the in vivo optical measurements obtained immediately prior to animal euthanasia.

Fig. 6.

Quantification of NIR dye after extraction from tumor tissue samples at 48 h following exposure to sham or FUS therapy. Note that * and *** denote p-values <0.05 and 0.001, respectively.

IV. Discussions and Conclusion

In this study, we introduced different approaches for applying low-intensity multi-FUS therapy that were aimed at improving systemic drug delivery to cancerous tissue. These strategies were based on spatial and temporal excitation sequences and use of numerous US focal zones, which were termed multi-FUS-SE and multi-FUS-TSE. Experimental results from in vivo optical imaging of a fluorescent tracer (i.e. surrogate drug) injected intravenously and the subsequent ex vivo NIR dye extraction from excised tissue showed accumulation in the targeted tumor space was considerably improved during the multi-FUS-TSE therapy approach. This study was performed using a concentric US transducer, which allows beamsteering and targeted treatment of an extended user selected tumor region. A co-registered linear array transducer provides US image guidance and anatomical navigation. A recent study also reported use of a concentric array transducer and electronic beamsteering for positioning and placement of numerous US focal zones during tissue harmonic motion imaging [26]. This study showed that the use of electronic beamsteering increased the imaging area without physical movement of the transducer and shortened the total US imaging duration.

In another recent study, Lu et al. [31] combined dual frequency and split-focus approaches to substantially enhance tissue heating during high-intensity FUS (HIFU) treatment. Using a prototype broadband split-focus transducer, a number of factors impacting thermal ablation and MB cavitation were compared including single versus multi-FUS strategies. During thermal ablation, they observed a 6-fold increase in the lesion volume when using a multiple FUS approach at twice the input power of the single FUS case. Later, this concept was used in a multi-focal HIFU method to achieve mild heating and cavitation reduction [32]. In addition to this, recent clinical studies used a multi-FUS sequence to increase microvascular permeabilization in volume space [27]–[29]. These studies used a power ramp test to find the inertial cavitation threshold, which is specific to each application. However, there have been no studies to date that have reported the use of a multi-focal zone positioning strategy applied over an extended tumor region to improve microvascular permeabilization and drug delivery. In short, the specialized multi-FUS therapy system used in the current work and presented herein is a promising novel approach for improved cancer treatment.

Many studies have investigated the relationship between various US and MB parameters that have been used for different applications like increased microvascular permeabilization and drug delivery. Typical peak negative pressure values that produce stable MB cavitation are in the range of 0.3 to 2.3 MPa for a 2.0 MHz US pulse [8], [9], [27]. It has also been shown that the inertial MB cavitation threshold for inducing microvascular permeabilization increases as the square root of the therapeutic US frequency [8]. This aside, use of lower frequencies is typically preferred during therapeutic US studies since these lower frequencies are less aberrated by tissue and bone structures during transmission. In vitro studies have also revealed that the maximum uptake of extracellular molecules in cancer cells occurred at an US frequency of 1.0 MHz and MI of 0.5 [9], [33]–[37]. Notwithstanding, tissue damage (bioeffects) from use of FUS exposure can occur and is due in part to the physical properties of the MB, US frequency of exposure, and duration of inertial MB cavitation. Intimately related to the latter, it has been suggested that the MI value is a key parameter that helps differentiate the onset of stable versus inertial MB cavitation during application of therapeutic US [8]. The chosen MI used in this study (i.e. MI < 0.5) is known to produce stable cavitation. The nature of MBs (half-life, size, dose, and exposure time) have a large effect on the US therapeutic procedure. In this study, 2.3 × 10⁷ MBs (1 × 10⁶ MBs/mL) were administered as a slow bolus injection (half-life in small animals of about 5 min). Per the manufacturer package insert, these MB populations have a mean diameter in the range of 1.1 to 3.3 μm with 98% less than 10 μm.

Our preliminary data presented herein revealed some interesting findings. Experimental results indicated that multi-FUS-TSE therapy produced a larger accumulation of the fluorescent tracer in the targeted tumor space compared to single-FUS and multi-FUS-SE US therapies. However, these results were not statistically significant (p > 0.05). This could be due to small sample size and/or due to the US parameters not being fully optimized for the multi-FUS sequence. Using the same US therapy settings, it has been shown that a single-FUS approach can produce a 10 to 15% increase in fluorescent dye accumulation within breast cancer-bearing mice [9]. During our current study, a 11.3% increase in fluorescent signal was found when using the single-FUS therapy compared to sham US therapy measurements.

The multi-FUS-SE therapy approach produced less NIR dye accumulation than the multi-FUS-TSE therapy. This might be counter intuitive, since the multi-FUS-SE therapy could leave more time for MB replenishment at each focus between US pulses. However, more detailed studies have to be performed to draw additional conclusions from these results. Another limitation of this study is that we have not performed any measurement to track the thermal and mechanical bioeffects during experiments, which can directly influence in vivo results. Studies detailed in the literature have shown that the occurrence of mechanical bioeffects are more probable with short pulses at a low repetition rate of US exposure, whereas thermal effects are more likely when operating in a more continuous US mode [38], [39]. Temperature changes in a phantom material exposed to pulsed US energy were shown herein to be negligible in support of this statement. However, it is important to note that these phantom measurements were not performed in the presence of MBs, which can induce temperature increases during in vivo studies when combined with FUS [40]. Even though our experimental settings are within the threshold of inertial cavitation [38] and below those that would produce tissue heating [39], a feedback control mechanism will help to deliver and monitor the therapeutic US procedure. In general, therapeutic US parameters need to be optimized for the multi-FUS sequence as it is influenced by both the spatial and temporal aspects of the US focusing approach. US therapy was also affected by tumor size. Larger tumors can exhibit core necrosis and vascularization that is more dense in the tumor periphery. We believe to have mitigated the impact of these conditions by performing experiments on animals with mean tumor size comparable between groups.

The number of focal points used for the multi-FUS sequences ranged from 29 to 55 (40.7 ± 7.0), which is about 25% of the entire tumor area. From the results, we can conclude that the increase in tumor treatment area had a direct influence on accumulation of the fluorescent dye. This study helped evaluate multi-FUS strategies within a single plane and these strategies need to be extended to the whole tumor volume to maximize the therapeutic US effect. This necessitates development of mathematical models and simulations that will simplify and help determine the impact of changing US protocol parameters. The mathematical models can be developed using analytical expressions derived from passive cavitation images [41], [42]. Numerical studies should help find an optimal in vivo multi-FUS sequence for maximal anticancer treatment.

Another important dimension to US therapy is monitoring the drug delivery process, which can improve site specific drug deposition. Magnetic resonance-guided FUS (MRgFUS) drug delivery systems are emerging as a promising technology and are commercially available for therapeutic US applications. Even though MRgFUS systems have inherent advantages, they also have known limitations. First, they are not readily available due to the high cost of the systems and housing combined with the labor-intensiveness during operation and procedures. Where supported by the application, many researchers have worked on development of US image-guided FUS therapy systems [24] [25]. Passive acoustic mapping (PAM) is a promising imaging method for real-time monitoring of drug delivery in tumors by mapping the MB cavitation process [43]. In a different study, the same transducer was used for both therapy and monitoring simultaneously, where the received signal was processed to separate cavitation from the therapeutic US signal [44]. The US image-guided FUS system used for our in vivo experiments can also be integrated with real-time therapeutic monitoring and control, and this advance constitutes future work.

Biographies

Lokesh Basavarajappa Lokesh Basavarajappa working as a postdoctoral researcher in the Department of Bioengineering at the University of Texas at Dallas, USA. He received Bachelor and Master degrees from Visvesvaraya Technological University, Belgaum, India. He then received his Ph.D. degree in applied mechanics (Biomedical engineering group), Indian Institute of Technology Madras, Chennai, India. His research interests include biomedical ultrasound, quantitative ultrasound, ultrasound elastography, ultrasound therapy, and photoacoustic imaging.

Girdhari Rijal Girdhari Rijal is an assistant Professor in the Department of Medical Laboratory Science and Public Health at Tarleton State University (TSU). He received B.S. degree in both general biological sciences and Medical Laboratory Technology followed by M.S. in Medical Microbiology. He graduated with Ph.D. degree in Regenerative Medicine and Pathology from School of Dentistry, Kyungpook National University, South Korea. He did postdoctal fellowships from POSTECH, South Korea and Washington State University, USA. He worked as a Research Scientist and the Lab Manager at department of Bioengineering at University of Texas before joining TSU.

Kenneth Hoyt Kenneth Hoyt is an Associate Professor in the Department of Bioengineering at the University of Texas and Dallas and Department of Radiology at the University of Texas Southwestern Medical Center. He has been an IEEE Member since 1999. He received a B.S. degree in Electrical Engineering from Drexel University (Philadelphia, PA) in 2001, followed by M.S. and Ph.D. degrees in Biomedical Engineering in 2004 and 2005, respectively, from the same institution. He did a postdoctoral fellowship in the Department of Electrical and Computer Engineering at the University of Rochester. Dr. Hoyt was faculty in the Department of Radiology at the University of Alabama at Birmingham (UAB) from 2008 to 2015. During this tenure he also received an M.B.A. degree from the School of Business (2011). Dr. Hoyt was elected fellow of the American Institute of Ultrasound in Medicine (AIUM) in 2014. In short, Dr. Hoyt’s research focuses on the development of novel ultrasound imaging strategies for improved human disease management (e.g., cancer and diabetes).