Immunomodulation and targeted drug delivery with high intensity focused ultrasound (HIFU): Principles and mechanisms.

HIFU is of interest to physicians and research scientists because it provides a non-invasive paradigm for treatment of a vast range of diseases for which the conventional therapies have major limitations. With advancement in technologies such as phased array transducers, real-time feedback, and temperature monitoring with ultrasound guidance or MR thermometry, new and exciting applications for HIFU have evolved. Currently, HIFU is FDA approved for the treatment of uterine fibroids, breast tumors, bone metastases, prostate cancer, and neurological applications such as essential tremor and Parkinson’s disease ( Elhelf et al., 2018 ; Izadifar, Izadifar, Chapman, & Babyn, 2020 ; She, Cheung, Jenkins, & Irwin, 2016 ). Focused ultrasound is a safe, non-invasive stimulus for controlled drug release from nanocarriers that are sensitive to elevated temperatures or pressures ( Couture, Foley, Kassell, Larrat, & Aubry, 2014 ). Therefore, although the encapsulated drug is circulating through the body, it is only released in the region where HIFU is applied. In this way, we can achieve triggered, on demand drug release in the diseased area with zero to minimum side effects systemically. HIFU-induced hyperthermia ( Frazier et al., 2016 ; May & Li, 2013 ) and cavitation ( Liang et al., 2010 ) have been shown to augment NP accumulation and increase local drug bioavailability especially from TSLs. More groups are reporting data on the use of HIFU in conjunction with immunomodulation drugs and/or chemotherapeutic drugs to boost the antitumor responses in various tumor types. In addition to tumor ablation, application of HIFU has advanced to include targeted drug ( McClure, 2016 ) and gene delivery ( Rahim et al., 2006 ), opening the blood-brain barrier ( Hynynen, McDannold, Vykhodtseva, & Jolesz, 2001 ), and thrombolysis ( Pajek, Burgess, Huang, & Hynynen, 2014 ). Currently, investigations with HIFU are ongoing for treatment of polycystic ovarian syndrome and atherosclerosis ( Shehata et al., 2013 ; Shehata et al., 2014 ). Additionally, HIFU clinical trials are ongoing for the treatment of soft tissue tumors and osteoid osteoma, with promising results ( Ghanouni et al., 2017 ; Sharma et al., 2017 ). Tables 6 and 7 list the current active ongoing HIFU human and veterinary clinical trials, respectively. A significant amount of recent scientific work has been devoted to optimizing the various types of drug delivery systems, including microbubbles ( Han, Lee, Kim, & Kim, 2017 ), nanobubbles ( Batchelor et al., 2020 ; Yao et al., 2016 ), liposomes ( Haemmerich & Motamarry, 2018 ; Zangabad et al., 2018 ), and polymeric NPs ( Yildirim, Blum, & Goodwin, 2019 ; Zeng, Liu, & Peng, 2021 ), so that they are ultrasound responsive. Clinical applications include anticancer therapy, cardiovascular system treatment, induction of an immune response, transdermal drug delivery, thrombolysis, disruption of the blood-brain barrier, and antibacterial applications ( Kasimanickam, Ranjan, Asokan, Kasimanickam, & Kastelic, 2013 ; Zhao, Du, Lu, Jin, & Ge, 2013 ). Future directions of HIFU research can focus on strengthening the combinatorial therapies especially with immunomodulatory NP delivery systems for enhanced anti-tumor immunity and targeting multiple immune-resistant pathways. Understanding the specific effects of HIFU against certain tumor types and why they are more immunomodulatory relative to other cancer sub-types, and how they can be overcome using an array of combinatorial regimens can enhance the pace of clinical translation. For example, mechanistically bridging HIFU with chemo- or radio-resistant tumors can generate a “team” of therapies that work in highly challenging situations of failed individual therapies, thereby synergistically augmenting response rates and overall patient survival. We hypothesize that HIFU will expand much more in the immunotherapy arena in the foreseeable future with various NPs. It will also be extended in the realm of laser and photon radiotherapy, and immune checkpoint inhibitor efficacy enhancement against poorly immunogenic tumor that are currently challenging to treat using conventional approaches. Performing high-intensity focused ultrasound (HIFU) in calcified regions, such as bones, is also particularly challenging due to sound wave attenuation. Therefore, new approaches and devices will need to be developed to enhance its application in such tissues, particularly by exploring novel thermometry methods. Furthermore, additional investigations are required to understand HIFU’s ability to reprogram the bone tumor microenvironment. This research is particularly crucial for understanding the feasibility of using HIFU against osteosarcoma, which is typically resistant to immune therapy. If HIFU can reprogram these tissues, it could significantly enhance outcomes, especially in pediatric patients. Similarly, while the application of HIFU for opening the blood-brain barrier continues to increase, it is unclear whether this approach could enhance immune checkpoint therapies in clinical settings. Future clinical trials incorporating design principles that combine HIFU with immune checkpoint therapies could provide novel insights into the potential of this combinatorial approach. This research could also be extended to the realm of combining HIFU with radiation therapy to reduce radioresistance, a topic which has received scant attention in clinical studies.

Application of HIFU can have diverse bio-effects. Below we describe the types of effects, which can be broadly categorized as tissue destruction, targeted drug delivery, immunomodulation, or other effects. Tissue destruction is the most common application of HIFU ( Al-Bataineh, Jenne, & Huber, 2012 ). HIFU-induced tissue destruction can be achieved by using thermal energy to induce protein denaturation ( Webb, Lubner, & Hinshaw, 2011 ) or by using mechanical energy to induce cell membrane destruction via cell stress ( Jang, Lee, Lee, Kim, & Hwang, 2010 ). The thermal effects of ultrasound for therapeutic purposes have been extensively explored for non-invasive treatment of a variety of clinical conditions ( Lynn et al., 1942 ) ( Kennedy, 2005 ), including tumors in the breast, and liver ( Elhelf et al., 2018 ), prostate ( Uchida et al., 2012 ), uterine fibroids ( Dobrotwir & Pun, 2012 ; Hesley, Gorny,& Woodrum, 2013), musculoskeletal pain ( Scipione et al., 2018 ); and neurologic disorders such as essential tremor, parkinson’s disease, and neuropathic pain ( Elias et al., 2013 ; Jeanmonod et al., 2012 ; Magara et al., 2014 ). During ablative treatment, high intensities of acoustic energy are deposited in the focal region by HIFU, which leads to rapid heating. Temperatures quickly rise to more than 55 to 60 °C within seconds, which causes rapid cellular necrosis in the targeted area with minimal damage to the surrounding tissue ( Fig. 2 ) ( Ter Haar, 2007 ). These exposures are usually high power, short duration exposures that produce well-demarcated coagulated lesions surrounded by apoptotic viable cells on the order of a few cellular layers ( Ter Haar & Robertson, 1993 ) ( Vykhodtseva, Mcdannold, Martin, Bronson, & Hynynen, 2001 ). Although lesion size depends on the equipment and parameters used, the classic individual thermal lesion can be the size of a rice grain ( Elhelf et al., 2018 ), which allows for extremely localized treatment and sharp treatment borders. The threshold for thermal necrosis depends on the temperature reached in the tissue, the rate of HIFU application, and the thermal sensitivity of the target tissue ( Nikfarjam, Muralidharan, & Christophi, 2005 ). Magnetic resonance (MR)-guided imaging allows for the monitoring of temperature rise in real time, allowing quantification of the therapeutic dose ( Jolesz, Hynynen, McDannold, & Tempany, 2005 ). Additionally, ultrasound imaging and tissue characterization techniques can be used for treatment monitoring for many clinical applications ( Song, Yoo, Song, & Chang, 2013 ). Using high power and very short pulse durations creates a large pressure change, low energy deposition, and minimal thermal rise in tissues. However, the high pressure exposure conditions lead to acoustic cavitation ( Leighton, 1994 ), which is the most prominent mechanical effect of HIFU. In general, acoustic cavitation can be thought of as the formation of a bubble cloud that is created as the traveling ultrasound waves interact with dissolved gases in the tissues in response to an acoustic pressure field. This phenomenon can occur during ultrasound propagation in water ( Atchley et al., 1988 ) as well as in biological tissues. The threshold of pressure amplitude that is required to initiate thermally significant cavitation in vivo is inversely proportional to the frequency of ultrasound treatment ( Hynynen, 1991 ). The nucleation of bubbles within a tissue can also be achieved by injecting microbubbles, apart from being generated by the ultrasonic peak negative pressure itself ( Chen, Hsu, Chung, & Yu, 2009 ; Deng, 2010 ). The pressure threshold for cavitation to occur depends on the acoustic parameters, such as frequency, pulse length, and peak pressure, as well as environmental properties such as the tissue type and tissue temperature ( Santos, 2020 ). Types of acoustic cavitation include stable and inertial cavitation ( Fig. 3 ) ( Shehata, 2012 ). Stable cavitation occurs when bubbles oscillate steadily and grow in size as the pressure changes at the focal point, creating shear stress and microstreaming of the surrounding fluid ( Elhelf et al., 2018 ). It can induce moderate changes at the cellular level, such as enhancing cell membrane permeability to drugs and other molecules ( Liang, Tang, & Halliwell, 2010 ). The main application of stable cavitation is to alter vascular permeability for increased drug/gene/nanoparticle (NP) extravasation, penetration, and thus improved delivery to whole tissues ( Yeh, 2012 ). By altering cell permeability and action potential via ion channel and receptor stimulation, it also facilitates drug delivery to cells ( Yang, Li, Du, & Chen, 2019 ). Inertial cavitation occurs when a certain pressure threshold has been exceeded, which causes a violent collapse of the microbubbles, which in turn releases a shockwave that is capable of destroying cell membranes and even liquefying or ablating cells at the focal point ( Chen et al., 2009 ; Wrenn et al., 2012 ). The use of inertial cavitation to create mechanical lesions with clearly demarcated margins of tissue fractionation is known as histotripsy ( Sukovich et al., 2018 ). By using injected microbubbles, which are easily visible with ultrasound imaging, to lower the threshold for inertial cavitation only at the target, acoustic cavitation can be monitored in real time to guide accurate targeting with the goal of minimizing unwanted tissue damage ( Kwan, Graham, & Coussios, 2023 ; Hempel et al., 2011 ; Wrenn et al., 2012 ). High precision and tissue selectivity are thus the well-established hallmarks of the histotripsy procedure ( Lin et al., 2014 ; Vlaisavljevich et al., 2013 ; Vlaisavljevich et al., 2015 ). Other physical phenomena such as acoustic streaming and radiation forces also contribute to mechanical-based bio-effects ( Dalecki, 2004 ). Unlike most other ablation therapies, histotripsy is non-thermal in most cases. Therefore, it is not affected by the heat-sink effect and remains safe and efficacious for use near tissues such as vasculature ( Pillai et al., 2015 ). Histotripsy-treated tissues have shown faster dissolution after treatment compared to the other ablation modalities ( Smolock et al., 2018 ; Vlaisavljevich et al., 2016 ). For example, healthy rat livers treated with cavitation-cloud histotripsy showed rapid shrinkage of treated volumes, granulation, and growth of healthy hepatocytes within 28 days post-treatment, with minimal scarring ( Vlaisavljevich, Greve, et al., 2016 ). Histotripsy has been investigated for a wide range of preclinical applications in large and small animal models for tumors, including liver cancer ( Vlaisavljevich et al., 2017 ; Vlaisavljevich, Greve, et al., 2016 ), prostate cancer ( Hempel et al., 2011 ), renal cancer ( Lake, Xu, Wilkinson, Cain, & Roberts, 2008 ; Longo et al., 2019 ), pancreatic cancer ( Allen & Vlaisavljevich, 2021 ; Hendricks, Brock, Gannon, Zeher, & Sereno, 2020), and musculoskeletal tumors ( Arnold et al., 2021 ), as well as for brain applications ( Ammi et al., 2008 ; Sukovich et al., 2016 ; Sukovich et al., 2018 ). Clinically, three Phase I trials have been conducted to investigate the safety of histotripsy in patients with liver cancer ( Vidal-Jove, 2023 ), calcified aortic stenosis ( Messas et al., 2021 ), and benign prostatic hyperplasia ( Kaplan, 2019 ). In particular, histotripsy of liver cancer patient tumors was found to be safe to administer, and interestingly also induced the shrinkage of non-targeted, untreated metastases likely via induction of anti-tumor immunity ( Vidal-Jove, Serres-Creixams, Ziemlewicz, & Cannata, 2021 ). Similarly, benign prostatic hyperplasia patients receiving histotripsy treatments from the perineum region showed an improvement in prostate symptom score with no adverse effects. Other non-tumor applications of histotripsy to calcified aortic valves using short, high-pressure ultrasound pulses likewise didn’t exacerbate the signs of stroke, myocardial infarction, transient ischemic attack, cognitive impairment, or treatment-induced death in patients; thereby indicate that histotripsy is a clinically safe non-invasive treatment modality. HIFU-assisted delivery of drugs/genes has gained increasing attention in recent years because it permits spatially confined delivery of therapeutic compounds in target areas, such as in tumors ( Zhu & Torchilin, 2013 ). Different HIFU mechanisms can be used for enhanced cellular drug uptake via sonoporation, vasodilation and increased vascular permeability, and drug delivery vehicles ( Yang et al., 2019 ). Sonoporation is the process whereby the pore size in the cell membrane increases because of ultrasound-induced mechanical impact via stable cavitation. Stable cavitation physically disrupts the integrity of membrane assembly, leading to membrane poration ( Liang et al., 2010 ). The reversible changes formed at the cellular level allow passive entry of drug molecules, genes, or NPs into cells ( Liang et al., 2010 ). Additionally, stable cavitation produces microstreaming, which increases the flow of fluid in a cell’s environment. This increase in flow may further assist the opening of pores, allowing a greater volume of drugs/compounds to be directed towards the cells, which enhances cellular uptake ( Collis et al., 2010 ). Ultrasound-induced sonoporation for improved delivery of DNA to mammalian cells was first demonstrated by Fechheimer et al. in 1986 ( Fechheimer, Denny, Murphy, & Taylor, 1986 ). This method can also be used in vivo for delivery of genetic material, which greatly increases the specificity of treatments compared to the alternatives ( Greenleaf, Bolander, Sarkar, Goldring, & Greenleaf, 1998 ). Focused ultrasound-assisted gene therapy can be used to treat cancers like glioblastoma and squamous cell carcinoma, cardiovascular diseases and Parkinson’s ( Han et al., 2020 ; Huber et al., 2003 ; Newman & Bettinger, 2007 ; Yang et al., 2021a ; Ziadloo, Xie, & Frenkel, 2013a ). Ziadloo et al. combined local tumor necrosis factor-alpha (TNF-α) plasmid gene therapy with pulsed-focused ultrasound (pFUS) in a murine squamous cell carcinoma tumor model. pFUS exposure at 1 MHz using 50 ms pulses and a pulse repetition frequency of 1 Hz given before intratumoral injection of TNF-α achieved relatively superior tumor regression compared to monotherapies ( Ziadloo, Xie, & Frenkel, 2013b ). Huber et al. similarly achieved enhanced transfection of intratumorally injected plasmid DNA in a rat dunning prostate tumor model with sinusoidal continuous wave ultrasound consisting of several bursts ( Huber & Pfisterer, 2000 ). Notably, their data suggested the presence of relatively higher populations of ultrasound-transfected cells in vitro than in the rat model due to the high wash rate of DNA by blood flow, poor intratumoral distribution, and small US focal spots (3 mm) compared to whole tumors (∼10 mm). Physiological barriers between the blood vessels and their surrounding tissues can limit delivery of drugs to the intended targets. HIFU can reversibly cause widening of blood vessels (vasodilation) and increase their permeability, thereby increasing blood flow and temporarily allowing drugs to pass through them and into the targeted region. HIFU can create a pressure change at the applied location, triggering the endothelium of targeted blood vessels to release nitric oxide, the chemical signal that causes smooth muscle relaxation and the dilation of blood vessels ( Maruo et al., 2004 ). Vasodilation can be induced in ischemic tissue to enhance the effects of radiotherapy by increasing the delivery of oxygen and blood to the target. Vasodilation can also aid treatments by increasing the amount of the drug delivered into the targeted tissue/region. Kang et al. elegantly extended this idea in the context of nanomedicine by complexing a HIFU responsive N-heterocyclic carbene-based nitric oxide donor micelles with doxorubicin loaded micelle to increase the on-demand nitric oxide release, vasodilation, and drug accumulation at the tumor site ( Kang et al., 2019 ). It is a reversible process, with no permanent damage to targeted tissue since pulsed HIFU causes minimal thermal effects ( Yang, Chiu, Liu, Lin, & Ho, 2009 ). Chen et al. showed significantly increased delivery of doxorubicin by ∼60% to prostrate tumor compared to Dox only group with pulsed focused ultrasound (pFUS) exposures ( Chen, Cvetkovic, Ma, & Chen, 2012 ). The delivery of drugs across vessel walls is controlled by a network of endothelial cells joined by tight junctions. The mechanical effects of HIFU disrupt these tight junctions to increase permeability, which can be more efficiently induced by using microbubbles ( Sheikov, McDannold, Sharma, & Hynynen, 2008 ). This same effect had been employed to non-invasively open the blood-brain barrier, which has unlocked a vast array of potential treatments for many neurological disorders, including Parkinson’s disease, Alzheimer’s disease, and glioblastoma ( Fan et al., 2015 ; Samiotaki, Acosta, Wang, & Konofagou, 2015 ; Tsai, 2015 ). Sheikov et al. provided the first direct evidence of the disruption of tight junction-specific transmembrane proteins claudin-1, claudin-5, occludin, and zona occludens (ZO-1) with focused ultrasound at a set frequency of 1.5-MHz in combination with microbubble-based contrast agent (Optison). The loss of the junctional barrier could be observed up to 4 h, thereby allowing the paracellular transport of compounds or drugs through the blood-brain barrier, highlighting the value of the combinatorial approach ( Sheikov et al., 2008 ). Yang et al. developed a lipid-polymer hybrid nanoparticles (NPs) to deliver CRISPR/Cas9 plasmids targeting O6-methylguanine-DNA methyltransferase (MGMT), a gene that regulates temozolomide (TMZ) mediated DNA damage, in murine glioblastoma tumors using ultrasound. Compared to NP alone, FUS microbubble combined NP gene therapy demonstrated superior tumor growth and prolonged survival of tumor-bearing mice without any evidence of adverse effects ( Yang et al., 2021b ). Mooney el al. administered a microbubble contrast agent (Definity) and delivered focused ultrasound immediately to the rat hippocampus region in 10 ms bursts at a 1 Hz burst repetition frequency for 120 s to enhance delivery of an antidepressant. Although a long-term cessation of depressive symptoms was not achieved, data suggested a short-term antidepressant effect ( Mooney, Nobrega, Levitt, & Hynynen, 2018 ). MR-guided focused ultrasound (MRgFUS) with intravenous microbubble is also getting translated to phase I clinical trial against neurodegenerative diseases. Gasca-Salas et al. conducted a prospective, single-arm, non-randomized, proof-of-concept, safety, and feasibility study in five patients with Parkinson’s disease with dementia (PDD). The combinatorial exposures opened the parietooccipito-temporal junction in 8 out of 10 treatments, which was associated with mild cognitive improvements ( Gasca-Salas et al., 2021 ). These are promising advancements, however, additional large multi-site clinical trials with long-term follow-ups are needed to establish the feasibility and reliability of this approach. In contrast to mechanical approaches, hyperthermia is another way to induce vessel changes. Hyperthermia heats tissues to 40–45 °C for several minutes to increase blood flow ( Thanou & Gedroyc, 2013 ), thereby drastically increasing tissue perfusion ( Pawar & Joshi, 2013 ), and oxygen delivery to the area, thus enhancing the metabolic activity and sensitivity of the targeted cells to drugs, especially tumors ( Ko & Dollinger, 2008 ). Because HIFU has the ability to penetrate into the deeper tissues of the body with precision, it has numerous and wide-ranging potential clinical applications ( Foundation FU, 2015 ). Tissue temperatures can be monitored in real time using MR imaging or interstitial temperature sensors (e.g., thermocouples), which allows for accurate temperature control. This approach has been applied in patients since the 1970s ( Marmor, Pounds, Postic, & Hahn, 1979 ). Some of the earlier patient treatments with focused ultrasound for hyperthermia used invasive thermocouples to monitor the temperature of heated tumors ( Shimm, Hynynen, Anhalt, Roemer, & Cassady, 1988 ). However, use of invasive thermocouples can introduce errors when placed within the focus of an ultrasound field ( Hynynen, Martin, Watmough, & Mallard, 1983 ), so non-invasive methods such as MR thermometry are currently preferred. Local hyperthermia has been reported to improve standard chemotherapy, especially with drug delivery systems based on liposomes as described below, for treatment of different diseases ( Partanen et al., 2012 ; VanOsdol et al., 2017a ; Wardlow et al., 2016 ). Due to the large surface area-to-volume ratio of as well as the flexibility in controlling their chemical and physical properties, NP is a powerful platform for enhancing traditional chemotherapy ( Kim, 2016 ). NP drug delivery systems protect the drug from rapid degradation or clearance, thereby enhancing drug concentrations in the target tissues. Consequently, lower doses of drug are required, and undesirable side effects of chemotherapy can be minimized ( Nevozhay, Kańska, Budzyńska, & Boratyński, 2007 ). This is especially important when there is a disparity between drug dose/concentration and the resultant therapeutic/toxic effects. Once the drug-loaded NPs reach the diseased tissues, the therapeutic agents are delivered. This release of drugs can be controlled through changes in the physiological environment, such as temperature, pH, osmolarity, or enzymatic activity ( Wilczewska, Niemirowicz, Markiewicz, & Car, 2012 ). Of these, the combination of HIFU with liposome-based NP carriers is clinically relevant. Liposomes were first discovered by Bangham and Horne in 1963 at the Babraham Institute in Cambridge, when they were testing the institute’s new electron microscope by adding negative stain to dry phospholipids ( Bangham, 1993 ; Kumar, Badde, Kamble, & Pokharkar, 2010 ). Liposomes are spherical bilayer phospholipid-based membranes that can be fabricated to be anywhere between a few nanometers to micron sized. Liposomes are biocompatible and can be broken down and integrated into cell walls. They are capable of holding drugs in their hydrophilic core or within their hydrophobic phospholipid bilayer coat. Liposomes provide a large drug payload per particle and protect the encapsulated drugs from metabolic processes ( Daraee, Etemadi, Kouhi, Alimirzalu, & Akbarzadeh, 2016 ). Improved drug delivery is realized by interactions of liposomes with cells and by enhancing drug solubility via membrane fusion, endocytosis, lipid transfer, and stable adsorption of liposomes into cells, thereby reducing toxicity to healthy cells, increasing the half-life of encapsulated drugs, and shielding them from the phagocytic system ( Wilczewska et al., 2012 ). The biophysical characteristics of NPs, such as vesicle size, lamellarity, surface charge, membrane fluidity, and surface, can be modified by the lipid composition and/or preparation method to confer various properties, including active targeting to substrates, temperature sensitivity, and immunological responses. Numerous liposomal drug formulations are in clinical use for treating different pathologies, such as anticancer, neurological, antibiotic, antifungal, anti-inflammatory, and antirheumatic drugs ( Antimisiaris, Marazioti, Kannavou, Natsaridis, & Gkartziou, 2021 ; Bulbake, Doppalapudi, Kommineni, & Khan, 2017 ; Mishra, Shandilya, & Mishra, 2018 ). In particular, temperature sensitive liposomes can be especially effective in combination with HIFU. The first formulation of liposomes to preferentially release their entrapped payload at transition temperatures was described by Yatvin et al. in 1978 ( Yatvin, Weinstein, Dennis, & Blumenthal, 1978 ). TSLs allow for external triggering of drug release both spatially and temporally. When the external temperature is raised above the melting phase transition temperature (T m ) of the lipids forming the bilayer, the structure of the bilayer changes as a transfer from a solid gel phase to a liquid-crystalline phase occurs ( Fig. 4 ). The release of hydrophilic drugs like doxorubicin HCl and ciprofloxacin HCl is highest at temperatures around the T m when the bilayer membrane is in the liquid-crystalline phase ( Kneidl, Peller, Winter, Lindner, & Hossann, 2014 ). This release generates high local drug concentrations and increased tissue penetration when used in combination with external stimuli such as HIFU heating (as explained in the previous sections)( Grüll & Langereis, 2012 ). Dipalmitoylphosphatidylcholine (DPPC) is the major component of most TSL formulations because its T m is above that of body temperature (i.e., 41.4 °C) ( Demel & De Kruyff, 1976 ; Mabrey & Sturtevant, 1976 ). Unwanted drug leakage at body temperature can be reduced by mixing DPPC with small amounts of other phospholipids, such as distearoylphosphatidylcholine (T m = 54.9 °C) ( Bassett, Anderson, & Tacker, 1986 ; Gaber, Hong, Huang, & Papahadjopoulos, 1995 ; Maruyama, Unezaki, Takahashi, & Iwatsuru, 1993 ). The overall T m of the formulation is determined by the composition of miscible phospholipids ( Hossann et al., 2007 ). Additionally, amphiphilic molecules (surfactants) are often used in liposomal formulations, but they potentially affect the vesicle stability to some extent ( Ickenstein, Arfvidsson, Needham, Mayer, & Edwards, 2003 ). For example, lipid-grafted PEG (e.g., 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N- [methoxy (polyethylene glycol)-2000], DSPE-PEG2000) is commonly used in liposomes to create a steric barrier for inhibition of uptake by the reticuloendothelial system as well as an increased blood circulation time ( Allen, Hansen, Martin, Redemann, & Yau-Young, 1991 ; Needham, McIntosh, & Lasic, 1992 ; Papahadjopoulos & Gabizon, 1990 ). Lyso PC (e.g., 1-stearoyl-2-hydroxy-sn-glycero-3-phosphocholine) is incorporated to mediate drug release by formation of lysolipid-stabilized membrane pores in TSLs ( Landon, Park, Needham, & Dewhirst, 2011 ; Mills & Needham, 2005 ). Stable drug retention at body temperature in the presence of blood components and a long in vivo half-life combined with a fast drug release rate around T m are the prerequisites for TSL formulations. In addition to phospholipid composition, the heat-triggered drug release depends to some degree on the drug molecule encapsulated, vesicle size, and the presence of serum components ( Gaber et al., 1995 ; Hossann et al., 2010 ; Hossann et al., 2012 ; Limmer et al., 2014 ; Ranjan et al., 2009 ). Some of the first FDA-approved liposomal formulations were Doxil ® (liposomal doxorubicin, Janssen) and DaunoXome ® (liposomal daunorubicin, Galen, Northern Ireland) for anticancer therapy and Abelcet ® (liposomal amphotericin B lipid complex, Sigma-tau) for fungal infections, and they were non-thermosensitive in nature ( Patra et al., 2018 ). Since their approval in 1995, several other liposomal formulations were approved for clinical use or are in clinical trials for the treatment of various medical disorders, including cancer, fungal infections, inflammatory conditions, neurological conditions, respiratory infections, infectious diseases, bacterial infections, and dermal conditions. The route of administration for these formulations varies depending on the active agent, indication, target tissue, and other factors and includes intravenous, intramuscular, intratumoral, ocular, oral, pulmonary, dermal, intranasal, intra-articular, and even vaginal (Antimisiaris, Marazioti, Kannavou, Natsaridis, & Gkartziou, 2021; Bulbake et al., 2017 ; Crommelin, van Hoogevest, & Storm, 2020 ; Kim & Jeong, 2021 ). Incorporation of lysolipids in the membrane bilayer led to the development of low TSLs (LTSLs) by Needham et al. in 2000 ( Needham & Dewhirst, 2001 ). This formulation is characterized by ultra-fast drug release upon heating, and approximately 70% of lysolipid was found to dissociate within 1 h post-intravascular injection ( Banno et al., 2010 ). The LTSL formulation with encapsulated doxorubicin (Thermodox ® ), licensed to Celsion Corporation (Columbia, MD, USA), is designed to be used in combination with heat-based treatments, such as radiofrequency ablation (RFA), microwave hyperthermia, or HIFU. It is currently under clinical investigation in combination with MR-HIFU for treatment of refractory/ recurrent solid tumors in humans as well as for sarcoma tumors in canine patients. Extensive research has been conducted and is underway to develop liposomal formulations for a variety of cancers ( Kneidl et al., 2014 ). Table 5 lists the immunologic and therapeutic effects of drug plus HIFU combinations for chemotherapy applications that are in preclinical and clinical studies. Local hyperthermia can improve standard chemotherapy including those that involves TSLs, for treatment of different diseases ( Partanen et al., 2012 ; VanOsdol et al., 2017a ; Wardlow et al., 2016 ). A key objective of this approach is to attain a stable temperature for a specific duration of time (tens of minutes to several hours) with a maximum volume of the target treated. Stable hyperthermia at 41 °C has been shown to be the most effective for drug deployment via LTSL systems ( Ponce, Vujaskovic, Yuan, Needham, & Dewhirst, 2006 ). Kong et al. also showed that there is a much higher probability that tumor blood vessels collapse above temperatures of 44 °C, which would result in a significant decrease in liposomes arriving via the microcirculation and hence a decrease in drug release in those regions. In one study, we combined clinical-grade Doxorubicin (Dox) encapsulated low temperature sensitive liposomes (LTSLs) with a clinical magnetic resonance-guided high intensity focused ultrasound (MR-HIFU) hyperthermia (40–41 °C) in rabbits with Vx2 tumors within superficial thigh muscle. This approach achieved significantly higher tumor doxorubicin concentrations (7.6- and 3.4-fold greater compared to free doxorubicin and LTSL ( Ranjan et al., 2012 ). We also found that nanobubble encapsulation in LTSLs improved Dox penetration in the tumor periphery and core following sequential HIFU hyperthermia, likely via enhanced opening of pores between endothelial cells ( VanOsdol et al., 2017b ). Several other reports have similarly shown that HIFU-induced hyperthermia can significantly enhance the delivery of anticancer treatments to targeted tumor sites in vivo ( Zhou, 2014 ). Beside tumor treatments, localized delivery of antimicrobials to treat challenging infections like biofilms and intracellular infections is a promising approach ( Gao, Chen, Zhang, Zhang, & Zhang, 2018 ; Rukavina & Vanić, 2016 ). Research from our lab has demonstrated the use of LTSLs for localized antimicrobial delivery in combination with HIFU hyperthermia in vivo . Additionally, Wardlow et al. showed that at higher temperatures (42 °C), extended treatment caused S. aureus bacterial membrane deformation and structural changes in the biofilm matrix, which consequently reduced bacterial viability in vitro ( Wardlow et al., 2016 ), and in a mouse chronic wound model ( Wardlow, Sahoo, Dugat, Malayer, & Ranjan, 2018 ). These results indicate strong promise for translation of LTSL-based systems with HIFU for wound treatment. The TSL plus HIFU vascular reprogramming can also integrate infectious disease and oncology. Kalyani et al. combined attenuated Salmonella strain with thermosensitive liposomes (thermobots) as a novel chemoimmunotherapeutic against colon cancer. Salmonella typhimurium facilitated the delivery of liposomes through chemotactic movement towards aspartate, ribose, and serine residues secreted by hypoxic tumor cells, thereby enhanced the localization of liposomes in the tumor region and HIFU triggered drug release ( Ektate et al., 2018 ). A healthy immune system can identify a broad range of pathogens and cancer cells. However, a compromised immune system leads to the advancement of chronic conditions such as cancer and chronic infections. Cancer cells or pathogens, adapt various mechanisms of immune evasion, like depletion of tumor-associated antigens in cancer or intracellular colonization of host/immune cells by pathogens ( Muthukrishnan, Masters, Daiss, & Schwarz, 2019 ), promoting the release of immune-suppressive cytokines ( Lindau, Gielen, Kroesen, Wesseling, & Adema, 2013 ). Extensive research has been conducted in the field of cancer immunology and its application—cancer immunotherapy—since the concept of cancer immunosurveillance was first proposed in 1957 by Burnet and Thomas ( Burnet, 1957 ). Antitumor immunity requires specific identification of cancer cells by the immune system of the patient for cancer elimination. To produce a cancer-specific response, the cancer cells need to have tumor-associated antigens. However, the majority of cancer types are immunologically unresponsive. HIFU is a noninvasive technique that can potentially turn immunologically cold tumors into immune-responsive hot tumors ( Eranki et al., 2020a ). When a tumor is ablated, it releases cellular proteins, debris, and neo-antigens that can trigger an immune response to the tumor at the primary site as well as potentially at distant metastases ( Finley et al., 2011 ; Haen, Pereira, Salih, Rammensee, & Gouttefangeas, 2011 ). Depending on the HIFU parameters used (thermal ablation, histotripsy (mechanical cavitation), or hyperthermia), there may be different immune responses as described below. Thermal ablation by HIFU has been shown to upregulate the expression of intracellular molecular chaperones (e.g. heat shock protein 70 (HSP70), in vitro and ex vitro ) ( Hundt, O’Connell-Rodwell, Bednarski, Steinbach, & Guccione, 2007 ; Kruse, Mackanos, O’Connell-Rodwell, Contag, & Ferrara, 2008a ), which can result in potent cellular immune responses. When a tumor is targeted with HIFU ablation, there is a marked increase in antigen presenting cells, such as dendritic cells. Table 2 lists and explains the immunologic and therapeutic effects of thermal ablation mediated tumor immunomodulation in preclinical and clinical settings ( Dromi et al., 2009 ). The increased abundance of tumor antigens released are recognized by the antigen presenting cells, which in turn activate lymphocytes for a specific immune response ( den Brok et al., 2004 ). Increased tumor infiltrating lymphocytes (TILs), particularly cytotoxic CD8+ cells and natural killer (NK) cells, along the ablated region can increase the chance of clearing surviving cancer cells ( Lu et al., 2009 ). Immunosuppression in a patient with a malignant tumor is a major obstacle in cancer treatment. In a study investigating changes in the circulating level of immunosuppressive cytokines in patients with malignancy before and after HIFU treatment, serum immunosuppressive cytokine levels decreased after HIFU. Significant decreases in vascular endothelial growth factor (VEGF), transforming growth factor (TGF)-β1, and TGF-β2 after HIFU treatment were observed in addition to direct tumor destruction ( Zhou et al., 2008 ). Clinical evidence suggests that HIFU treatment may also enhance local antitumor immunity in prostate cancer patients ( Kramer et al., 2004 ) and upregulate expression of HSP70 in breast cancer tumor debris ( Wu et al., 2007 ). Researchers have also described B cell activation, the maintenance of plasma cells, dendritic cell activation, and even the generation of tumor-specific antibodies following local tumor ablation ( Widenmeyer et al., 2011 ; Wissniowski et al., 2003 ). In a prospective study investigating the adaptive immune responses of radiofrequency ablation (RFA) against tumor-associated antigens (TAA) in patients after treating secondary liver tumors, Widenmeyer et al. found that only 6 of 49 patients studied had increased antibody levels or TAA-reactive CD8+ T cells months after treatment. Although the number of TAAs tested was limited, the study pointed to a weak effect of RFA on adaptive immunity ( Widenmeyer et al., 2011 ). Although many promising preclinical and clinical studies have shown that ablation therapies can control tumor growth locally and have positive effects on the systemic immune response, there have also been reports of the opposite. There are several preclinical reports of the RFA treatment of hepatocellular carcinoma tumors and colorectal metastases potentially stimulating growth of existing and distant tumors ( Ahmed et al., 2018 ; Ahmed, Kumar, Moussa, Rozenblum, & Goldberg, 2013 ). Studies investigating this issue have linked these pro-oncogenic effects to upregulation of inflammatory pathways in this region, including the IL-6-HGF/c-Met-STAT3-VEGF axis and the HSP70-related pathways ( Ahmed et al., 2016 ; Ahmed et al., 2018 ; Velez et al., 2016 ). Interestingly, Ahmed et al., demonstrated that a higher temperature/shorter duration treatment paradigm with adjuvant HSP inhibitors limited off-target systemic pro-oncogenic effects after hepatic RFA of rat livers ( Ahmed et al., 2018 ). Increased tumor growth after the minimally invasive ablation of tumors may be due to the tumor cell seeding that can occur along the needle track ( Distelmaier et al., 2017 ). This particular issue is avoided in HIFU ablation, which uses a completely non-invasive extracorporeal transducer system for highly focused ultrasound treatment application. Jenne et al. compared HIFU ablation with other thermal ablation methods and noted that controlling the variability of lesion shape and size was the major advantage of HIFU ( Jenne et al., 2007 ). Other advantages of HIFU are its ability to ablate tumors in difficult locations, such as close to the heart or diaphragm, where RFA would be technically impossible ( Cheung et al., 2012 ). Furthermore, the HIFU-associated heat-sink effect is minimal due to extracorporeal energy delivery, unlike RFA, in which energy is delivered using an active treatment electrode, which can lead to incomplete tissue ablation ( Cheung, Ma, & She, 2021 ). Histotripsy-induced cavitation breaks down tissues, ablating cells into subcellular fragments and acellular debris, which results in modulation of immune processes, cells, and molecules. Histotripsy’s postulated roles in immunomodulation are categorized as decreasing pro-tumor immune cells, cellular immunity, and systemic immunity ( Hendricks-Wenger, Hutchison, Vlaisavljevich, & Allen, 2021 ; Vlaisavljevich et al., 2016 ). Table 3 lists and explains the immunologic and therapeutic effects of histotripsy mediated tumor immunomodulation in various tumor models. Pahk et al. demonstrated that the supernatant from breast cells treated with boiling histotripsy (BH) can polarize THP-1 human monocyte cells to M1 macrophages and also repolarize M2 macrophages to the M1 state ( Pahk et al., 2019 ). Cytokine analysis revealed increased pro-inflammatory signaling molecules, including tumor necrosis factor (TNF), which is a potent and well-established M1 stimulating cytokine ( Kratochvill et al., 2015 ). In vivo , this would help to decrease the presence of pro-tumor immune cells and potentially alter the tumor microenvironment from cold to hot. Although not confirmed by the literature, this hypothesis is consistent with results from some of our preclinical studies. For example, we treated murine melanoma with histotripsy and found a strong correlation between improved immune effects and an increase in mouse survival ( Singh, Sethuraman, Miller, Malayer, & Ranjan, 2021a ). Histotripsy generates subcellular fragments through mechanical fractionation and induces sonic stress to enhance the expression of damage-associated molecular patterns (DAMPs), which enhance tumor inflammation, and antitumor immune effects ( Eranki et al., 2018 ; Eranki et al., 2020a ; Hu et al., 2007 ; Khokhlova et al., 2011 ). In addition, critical cytokines and chemokines have been found to be significantly altered, especially interferon (IFN)- γ, in multiple histotripsy studies ( Eranki et al., 2020a ; Pahk et al., 2019 ; Qu et al., 2020a ; Schade et al., 2019a ). For example, an in vivo murine study of neuroblastoma treated with histotripsy showed an approximately 2-fold increase in IFN-γ ( Eranki et al., 2020a ). While IFN- γ is the most consistently reported cytokine across therapies and tumor types, researchers have reported upregulation of other important cytokines, including interleukin (IL)-6, IL-2, TNF, IL-8, IL-13, and IL-10, post-histotripsy treatment ( Schade, 2023 ; Eranki et al., 2020a ; Pahk et al., 2019 ; Qu et al., 2020a ). In addition to cytokines, changes in levels of growth factors, including granulocyte-macrophage colony-stimulating factor and VEGF, have been reported post-histotripsy ( Eranki et al., 2020a ), adding to the potential for histotripsy to shift a cold tumor to a more proinflammatory and tumor-suppressive microenvironment. In response to DAMPs and the antitumor mediators (cytokines and chemokines), the cells associated with innate immunity are rapidly recruited. For histotripsy, this includes neutrophils, NK cells, dendritic cells, and macrophages ( Hu et al., 2005 ; Qu et al., 2020a ). The consequent modulation of adaptive immune cells has been strongly correlated with clinical success ( Keisari, 2017 ; Kepp, Marabelle, Zitvogel, & Kroemer, 2020 ). Reduction in the magnitude of T regulatory cells (Tregs) and increased ratio of CD8+ to Tregs in both the tumor-draining lymph nodes (TDLNs) and spleens of treated mice was reported ( Huang et al., 2012 ). Furthermore, re-challenge with tumor cells in the contralateral limb 6 days post-histotripsy showed reduced tumor growth, indicating treatment efficacy in generating systemic tumor-specific protection ( Hu et al., 2007 ). Notably, the effects of mechanical HIFU on minimizing the growth rates of treated tumors were more effective than thermal ablation, despite lesser debulking of primary tumor (43% and 85% tumor volume reduction, respectively). This suggests that the mechanical fractionation caused by HIFU stimulated a stronger immune response, and it may be employed in combination with thermal ablation to increase the efficacy of HIFU ( Hu et al., 2007 ). The increased release of DAMPs, altered inflammatory state, and enhanced cellular and systemic immune response caused by histotripsy can potentially improve the efficiency of immune checkpoint inhibitors (ICI) ( Eranki et al., 2020a ; Singh et al., 2021a ). Combining histotripsy with anti-CTLA-4 and anti-PDL1 was shown to induce potent systemic immunity and long-term immune memory to cure the majority of mice with unilateral and bilateral neuroblastoma tumors ( Eranki et al., 2020a ). In an ICI refractory melanoma model, the combination of histotripsy with anti-CD40 agonist antibody led to improved immune sensitization of tumors via the CXCL9-cytotoxic T cell axis, and consequently better mice survival rates, compared to the monotherapies ( Singh et al., 2021a ). HIFU-mediated mild-moderate hyperthermia has been shown to increase vascular permeability and cellular permeability and to enhance the metabolic activity of the hyperthermic targets. This enables enhanced delivery/passage of drugs and immune cells, increased bioavailability of drugs, and sensitivity to chemotherapy. While mild hyperthermia does not typically cause immediate cell death, it has been shown to alter the tumor microenvironment by inducing a myriad of cellular effects ( Tydings, Sharma, Kim, & Yarmolenko, 2020 ). As described in Table 4 , HIFU hyperthermia can directly promote antigen cross-presentation and tumor-specific T cell generation and expansion in a variety of tumor settings ( Baronzio et al., 2006 ). When cells are exposed to hyperthermia, they undergo a stress response. Viable tumor cells that evade the immune system experience an upregulation of HSP synthesis and surface expression. Consequently, HSP-peptide complex formation activates NK cells, which have direct antitumor activity. They also bind to dendritic cell receptors, facilitating their activation and consequent generation of cytotoxic T cells, augmenting direct antitumor activity. In addition to a local response, locally applied hyperthermia is also capable of increasing systemic levels of antitumor immune cells ( Baronzio, Gramaglia, & Fiorentini, 2006 ; Tydings, Sharma, Kim, & Yarmolenko, 2020 ). Wu et al. measured the levels of lymphocytes and NK cells in adults with solid tumors, including osteosarcoma, hepatocellular carcinoma, and renal cell carcinoma. Patients treated with HIFU showed an increase in CD4+ lymphocytes in their peripheral blood at 7–10 days post-treatment. CD4+ helper T cells play critical roles in initiating, regulating, and maintaining the antitumor immune response, especially the T cell-mediated antitumor response ( Wu et al., 2004 ). Local hyperthermia has the potential to enhance standard chemotherapy, especially when used with drug delivery system ( Partanen et al., 2012 ; VanOsdol et al., 2017a ; Wardlow et al., 2016 ). However, due to its ability to improve both local and systemic antitumor immune responses, hyperthermia is increasingly finding more applications in cancer immunotherapy. Singh et al. showed that a combination of HIFU hyperthermia and anti-CD-40 improved macrophage polarization dynamics and T cell functions to aid melanoma immunotherapy ( Singh et al., 2019 ). Sethuraman et al. assessed the combination of calreticulin (CRT) NPs with HIFU hyperthermia for immunogenic cell death-associated immunomodulation in a melanoma model. Their in vitro and in vivo data suggested enhanced antigen presentation by macrophages and infiltration of activated CD8+ T cells in tumors ( Sethuraman et al., 2020 ). The addition of HIFU hyperthermia led to modulation of the CRT-CD47-PDL1 axis, thus improving the overall therapeutic response against melanoma tumors.

HIFU acoustic intensities delivered to a well-defined focal region of tissue (focal zone) are generally 1000 to 10,000 times greater than those of traditional diagnostic devices ( Table 1 ) ( Ter Haar, 2007 ; Gt, 2001 ). This ability to ablate millimeter sized targets without harming surrounding tissues can position HIFU as an alternative treatment modality compared to radiation and surgery. Currently, treatment approaches based on HIFU for clinical indications or disorders are in various stages of development, with most in early stages of investigation ( Duc & Keserci, 2019 ). Worldwide, 34 indications have obtained regulatory approval for HIFU use. In the United States, HIFU has been approved by the FDA for seven indications (2021 State of the Field Report, Focused Ultrasound Foundation). To achieve localized effects, HIFU transducers are designed to converge beams of ultrasound waves into a single focal point. A transducer is a device that converts energy from one form to another. In HIFU, electrical energy is converted to sound energy using piezoelectric crystals that oscillate upon application of an alternating voltage, resulting in the generation of ultrasound waves in the receiving medium. Focusing is geometrically achieved using a spherical/concave curved transducer or by using a plane transducer with an acoustic lens to mimic a concave surface transducer ( Elhelf et al., 2018 ). Beam focusing can also be achieved via electronic steering in phased array transducers composed of multiple piezoelectric elements ( Fig. 1 ) ( Elhelf et al., 2018 ; Malietzis et al., 2013 ). The overall geometry of the transducer and frequency of the piezoelectric elements determine the characteristics of the transducer, including focal depth and width and length of the focal zone. An ultrasound wave from a transducer outside the body must travel through multiple tissue layers (including skin, subcutaneous fat, and muscle) before reaching the target organ. At each tissue interface, part of the energy carried by the sound wave is reflected, while the remaining energy is transmitted into the tissue. The acoustic impedance, Z, which is the product of the speed of sound and density of tissue, determines the transmission coefficient. Apart from fat, air, and bone, most human tissues have acoustic properties like those of water. Therefore, aqueous media are optimal for transmitting ultrasound energy from the transducer into the body. The loss in incident acoustic energy due to absorption/heating/scattering in a medium is characterized by its attenuation coefficient, which in turn is inversely related to the ultrasound frequency ( ter Haar & Coussios, 2007 ). An increasing attenuation is correlated with decreased water content and increased protein content. For example, tissues such as tendons, which have high protein content, have higher absorption of ultrasound waves and therefore greater heating effects compared to high water-containing tissues like blood ( Watson, 2008 ). Therefore, target tissue properties need to be considered when deciding on the HIFU treatment parameters.

Medical ultrasound is utilized for both diagnostic and therapeutic applications. Today, diagnostic ultrasound imaging is widely used for non-invasive examination of internal body structures (e.g., obstetric ultrasonography). However, the use of therapeutic ultrasound pre-dates diagnostic ultrasound ( Ter Haar, 2008 ). Therapeutic ultrasound generates a biological response by focusing acoustic energy in a defined region to produce thermal and non-thermal effects ( Nyborg, 2001 ). In the earliest reported investigation of focused ultrasound in 1942, Lynn et al. demonstrated biological effects in ex vivo liver samples and in vivo rodent models ( Lynn, Zwemer, Chick, & Miller, 1942 ). Despite its early start, exploration of the clinical applications of therapeutic ultrasound did not progress due to a lack of adequate image guidance to target and monitor treatments in real time ( Ter Haar, 2007 ). With recent advances in imaging, therapeutic ultrasound is making a resurgence in the medical community. One approach called high intensity focused ultrasound (HIFU) augments blood flow, stimulates tissue regeneration, or destroys tissue via thermal and mechanical effects ( Azhari, 2010 ). In this review, we describe the mechanisms by which HIFU elicits targeted physiological changes in tissue microenvironments and how it can be leveraged to improve drug delivery and immunotherapy in tissues.