Focused Ultrasound
US is a mechanical wave with a frequency above 20 kHz that is commonly used for medical, sonar, and industrial testing applications. Operating at frequencies above 1 MHz, medical imaging US creates images by interrogating tissue with mechanical waves. These mechanical or acoustic waves are generated and received by a US transducer, which converts an electrical input signal into an acoustic wave and vice versa. When these acoustic waves propagate through the body, the waves are reflected at tissue interfaces. Two common methods of transmitting and receiving these acoustic waves will be discussed in this review: pulse-echo imaging where the imaging transducer transmits acoustic waves into the body and then receives the reflected waves, and passive imaging where the imaging transducer only receives reflected waves generated by another transducer. When imaging in a traditional pulse-echo US system, the frequency of the transducer dictates the available imaging frequencies and therefore the axial resolution of the image. In linear arrays, the lateral resolution of the image can be limited by the transducer pitch, that is, the center-to-center element spacing. The geometry of the US transducer and its casing (collectively termed the probe) determine which regions of the body can be imaged or treated. The frequency of the wave depends on the requirements of the application, with high-resolution imaging calling for higher frequencies that have a smaller wavelength but greater attenuation over distance, that is, lower penetration depth. The relationship between the wavelength and the frequency is determined by the speed of the sound waves in the medium. Focused Ultrasound The US waves that are transmitted from the transducer, just like light, can be focused at a desired distance from the transducer by mechanical or electronic means. Mechanical focusing can be achieved by shaping the transducer (see Fig. A) or applying an acoustic lens to the transducer (see Fig. B). Shaping the transducer material (ie, the piezoelectric ceramic), the acoustic waves generated along its surface can be forced to converge at a single point (see Fig. A). An acoustic lens is a layer of varying thickness made of a material whose speed of sound is faster than in the interrogated tissue. The faster speed of sound at the edges of the lens allows the waves transmitted at those edges to reach the focus at the same time as the waves transmitted at the center of the lens (see Fig. B). Adjusting the radius of the curvature of the lens changes this focal depth according to Snell's law. Both methods of mechanical focusing allow for the use of low-cost, single channel electronics, but the focus cannot be adjusted in depth or steered from side to side. , Electronic focusing is achieved by transmitting waves from multiple acoustic sources at different starting times, allowing the waves to converge at a desired depth and to steer the focus from side to side (see Fig. C). This is the phased array approach that has played an important role in the adoption of FUS in the clinic. Phased arrays can both focus and steer the waves in real time by changing the delay between element transmissions. The tradeoff of this approach when compared with mechanically focusing is the added complexity and cost of the electronics and transducer. It is worth noting that mechanical and electronic focusing techniques can be combined to meet the specific needs of a given application. FUS provides therapeutic effects through different mechanisms of action. , There is a wide range of bioeffects in the tissue, and many mechanisms of action are still under investigation. , – A comprehensive list of FUS mechanisms of action is provided in the Focused Ultrasound Foundation State of the Field report. For the purposes of this review, we will briefly describe the mechanisms of action that benefit most from image guidance in current clinical practice and translational research. Thermal ablation uses HIFU to noninvasively heat tissue within the body beyond the threshold of protein denaturation, effectively acting as an ultrasonic scalpel to eliminate target tissue. Thermal ablation is used in the clinic to treat a range of disorders including essential tremor, prostate cancer, and uterine fibroids. Cavitation is the oscillation of microbubbles in the tissue. The expansion and contraction of these microbubbles, whether endogenous or exogenously introduced in the form of a US contrast agent, produce mechanical effects that can result in fracture, fragmentation, and tissue liquefaction. Histotripsy is a technique that takes advantage of these mechanical effects by applying short pulses with a large magnitude peak negative pressure (>5 MPa PNP). Drug delivery, including blood-brain barrier (BBB) permeabilization, also uses cavitation but at a lower pressure to induce changes in the cell membrane or the intercellular interfaces to enhance drug or other therapeutic delivery. These techniques typically use exogenous microbubbles. BBB opening is an example of this phenomenon, where the oscillation of the microbubbles causes the tight junctions of the BBB to temporarily open. , This permeabilization of the BBB is temporary and has been shown to improve delivery of a range of molecules with different sizes from small molecules to gene therapies through the usually impenetrable BBB. Still under clinical investigation, opening of the BBB has been shown to be safe in patients with Alzheimer’s disease, and many clinical trials are ongoing. , Neuromodulation is a burgeoning field within FUS. Low-intensity focused US waves are directed transcranially , , into the brain or to the peripheral nervous system outside of the skull to induce changes in the neural networks noninvasively at a high spatial precision without significantly heating tissue. Results in human subjects addressing depression, pain, and addiction show early promise. This field of study is in the very early stages and may benefit from real-time imaging techniques that can demonstrate what regions of the brain are affected when FUS is applied. Microbubbles Microbubbles are gas-filled, acoustically active micron sized spherical particles with a typical diameter of 1–10 μm. The gas core of the microbubble is surrounded by a shell, usually composed of phospholipids, proteins, sugars, or polymers. Commercial microbubbles available in the United States include Definity, Lumason, and Optison. These products are indicated for use as US imaging contrast agents. When a US wave acts on a microbubble, it causes the microbubble to expand and contract, oscillating and reflecting some of the US energy. Much of this reflected energy typically has a nonlinear relationship to the original US wave, providing a unique signature that allows the echoes from microbubbles to be distinguished from surrounding tissue which typically return echoes at the same frequency as the original wave. , In addition to their unique properties as US imaging contrast agents, microbubbles produce bioeffects that offer unique opportunities to enhance drug delivery (see below).   The US waves that are transmitted from the transducer, just like light, can be focused at a desired distance from the transducer by mechanical or electronic means. Mechanical focusing can be achieved by shaping the transducer (see Fig. A) or applying an acoustic lens to the transducer (see Fig. B). Shaping the transducer material (ie, the piezoelectric ceramic), the acoustic waves generated along its surface can be forced to converge at a single point (see Fig. A). An acoustic lens is a layer of varying thickness made of a material whose speed of sound is faster than in the interrogated tissue. The faster speed of sound at the edges of the lens allows the waves transmitted at those edges to reach the focus at the same time as the waves transmitted at the center of the lens (see Fig. B). Adjusting the radius of the curvature of the lens changes this focal depth according to Snell's law. Both methods of mechanical focusing allow for the use of low-cost, single channel electronics, but the focus cannot be adjusted in depth or steered from side to side. , Electronic focusing is achieved by transmitting waves from multiple acoustic sources at different starting times, allowing the waves to converge at a desired depth and to steer the focus from side to side (see Fig. C). This is the phased array approach that has played an important role in the adoption of FUS in the clinic. Phased arrays can both focus and steer the waves in real time by changing the delay between element transmissions. The tradeoff of this approach when compared with mechanically focusing is the added complexity and cost of the electronics and transducer. It is worth noting that mechanical and electronic focusing techniques can be combined to meet the specific needs of a given application. FUS provides therapeutic effects through different mechanisms of action. , There is a wide range of bioeffects in the tissue, and many mechanisms of action are still under investigation. , – A comprehensive list of FUS mechanisms of action is provided in the Focused Ultrasound Foundation State of the Field report. For the purposes of this review, we will briefly describe the mechanisms of action that benefit most from image guidance in current clinical practice and translational research. Thermal ablation uses HIFU to noninvasively heat tissue within the body beyond the threshold of protein denaturation, effectively acting as an ultrasonic scalpel to eliminate target tissue. Thermal ablation is used in the clinic to treat a range of disorders including essential tremor, prostate cancer, and uterine fibroids. Cavitation is the oscillation of microbubbles in the tissue. The expansion and contraction of these microbubbles, whether endogenous or exogenously introduced in the form of a US contrast agent, produce mechanical effects that can result in fracture, fragmentation, and tissue liquefaction. Histotripsy is a technique that takes advantage of these mechanical effects by applying short pulses with a large magnitude peak negative pressure (>5 MPa PNP). Drug delivery, including blood-brain barrier (BBB) permeabilization, also uses cavitation but at a lower pressure to induce changes in the cell membrane or the intercellular interfaces to enhance drug or other therapeutic delivery. These techniques typically use exogenous microbubbles. BBB opening is an example of this phenomenon, where the oscillation of the microbubbles causes the tight junctions of the BBB to temporarily open. , This permeabilization of the BBB is temporary and has been shown to improve delivery of a range of molecules with different sizes from small molecules to gene therapies through the usually impenetrable BBB. Still under clinical investigation, opening of the BBB has been shown to be safe in patients with Alzheimer’s disease, and many clinical trials are ongoing. , Neuromodulation is a burgeoning field within FUS. Low-intensity focused US waves are directed transcranially , , into the brain or to the peripheral nervous system outside of the skull to induce changes in the neural networks noninvasively at a high spatial precision without significantly heating tissue. Results in human subjects addressing depression, pain, and addiction show early promise. This field of study is in the very early stages and may benefit from real-time imaging techniques that can demonstrate what regions of the brain are affected when FUS is applied. Microbubbles Microbubbles are gas-filled, acoustically active micron sized spherical particles with a typical diameter of 1–10 μm. The gas core of the microbubble is surrounded by a shell, usually composed of phospholipids, proteins, sugars, or polymers. Commercial microbubbles available in the United States include Definity, Lumason, and Optison. These products are indicated for use as US imaging contrast agents. When a US wave acts on a microbubble, it causes the microbubble to expand and contract, oscillating and reflecting some of the US energy. Much of this reflected energy typically has a nonlinear relationship to the original US wave, providing a unique signature that allows the echoes from microbubbles to be distinguished from surrounding tissue which typically return echoes at the same frequency as the original wave. , In addition to their unique properties as US imaging contrast agents, microbubbles produce bioeffects that offer unique opportunities to enhance drug delivery (see below).   Microbubbles are gas-filled, acoustically active micron sized spherical particles with a typical diameter of 1–10 μm. The gas core of the microbubble is surrounded by a shell, usually composed of phospholipids, proteins, sugars, or polymers. Commercial microbubbles available in the United States include Definity, Lumason, and Optison. These products are indicated for use as US imaging contrast agents. When a US wave acts on a microbubble, it causes the microbubble to expand and contract, oscillating and reflecting some of the US energy. Much of this reflected energy typically has a nonlinear relationship to the original US wave, providing a unique signature that allows the echoes from microbubbles to be distinguished from surrounding tissue which typically return echoes at the same frequency as the original wave. , In addition to their unique properties as US imaging contrast agents, microbubbles produce bioeffects that offer unique opportunities to enhance drug delivery (see below).   No different than when diagnosing disease, different imaging modalities present tradeoffs when guiding therapy with FUS. This review will focus on magnetic resonance (MR) and US image guidance of FUS, but x-ray computed tomography (CT) and positron emission tomography/single photon emission computed tomography (PET/SPECT) imaging – also play an important role in FUS planning and posttreatment assessment. X-ray CT is a quick, readily available, and cost-effective method to measure tumor size and plays a key role in the RECIST (Response Evaluation Criteria in Solid Tumors) standard for assessing treatment effect on tumor burden. , PET/SPECT imaging provides opportunities to image immune and cell therapies in both clinical and preclinical studies. US and MR have been the most promising imaging modalities for FUS guidance to date due to their lack of ionizing radiation, their application for planning, monitoring, and posttreatment measurements, and their widespread availability in clinical settings (see Table ). Ultrasound Guidance Prior to the pairing of FUS with MRI performed by Hynynen in 1992, US was the dominant method for FUS guidance. , , In the context of FUS, US guidance presents the opportunity to image anatomy, image blood flow, image cavitation, measure a change in tissue properties, and monitor temperature changes. Though not all of these techniques are in clinical use, we will discuss each to provide insights into the benefits of US guidance and the direction of the field. Anatomical Imaging B-mode US guidance is used in the 3 indications most widely treated by FUS—uterine fibroids (36% of patients treated to date), liver tumors (21%), and prostate diseases (18%). By including a higher frequency imaging US transducer (>1 MHz) aligned with the low frequency FUS transducer (<1 MHz), a reliable detection of the FUS beam can be made in real time using pulse-echo imaging techniques. One model for imaging and treatment of the body is the table-based system, where the HIFU and imaging transducers are installed within a table with a water bath that is used to couple the transducers to the patient's body. With these platforms, used for uterine fibroids , and liver tumors, the treatment target can be selected using the B-mode US image as a guide. Then, as HIFU is applied, how well the tissue in the treatment site reflects the US waves (ie, echogenicity) can be monitored to determine if sufficient energy is being applied to cause ablation. A bright or hyperechoic spot, located at the focus of the HIFU beam, is the result of cavitation —the formation of gas microbubbles that then cause nonlinear reflections of the acoustic waves. In some cases, the B-mode image shows a hyperechoic bright spot at the site of tissue ablation after treatment, providing real-time imaging feedback of the ablation itself. , Although convenient, echogenicity alone is not a reliable method to determine the region of ablation or the extent of ablation. Prostate treatment uses a minimally invasive model, relying on a transrectal US probe that combines imaging and HIFU capability. The location of prostate cancer can be detected based on the hypoechoic (poorly reflected and appearing black on the US image) regions of interest, although this method of identifying tumors is highly operator dependent. , B-mode imaging is also used to guide histotripsy in the recently approved Histosonics Edison platform. Similar to HIFU systems used for thermal ablation, the Edison platform uses US guidance to plan and monitor treatment in real time. Following treatment with histotripsy, the tissue can exhibit a hypoechoic result (lacking US scattering), indicating liquefaction of the tissue. The detection of the cavitation cloud in real time is used to adjust the pressure amplitude of the applied US to induce the desired histotripsy effects (see Fig. ). , A similar approach is applied to select pressure amplitude in US-guided thermal ablation. Doppler Imaging—Blood Flow Doppler US imaging can be used to detect blood flow in real time. This offers the opportunity to monitor changes in tissue perfusion and activity that may allow detection of tumors as well as the effects of therapy on a tumor. – With FUS neuromodulation, Doppler US imaging has been proposed as a tool to monitor how the US modulates neuronal activity in real time. , Recent advances in high speed US imaging have opened up new opportunities to detect slow blood flow and higher-resolution imaging of blood flow—domains not previously able to be imaged with Doppler US. Contrast-Enhanced Ultrasound Contrast-enhanced US uses exogenous microbubbles to image the vasculature within the body. Due to their size, microbubbles can travel throughout the capillaries and other blood vessels of the body but do not pass into the interstitial spaces of the tissue. The nonlinear scattering of the microbubbles provides an opportunity to isolate their echo from that of the surrounding tissue to detect perfusion and blood vessel anatomy. The simplest techniques use the increased echogenicity where microbubbles are present to detect perfusion. Within the clinic, contrast-enhanced US has been used as an alternative to MRI to assess the nonperfused volume of uterine fibroids following ablation of the fibroid with HIFU. , Other opportunities for contrast-enhanced US techniques can be found in imaging other forms of cancer and the brain during BBB opening. Cavitation Detection Cavitation plays a significant role in multiple FUS treatments. To better characterize and control this mechanical bioeffect, different approaches have been adopted to selectively identify the cavitation activity of endogenous and exogenous microbubbles. This is possible due to the unique properties of the gas-filled microbubbles that scatter US waves. There are 2 domains of microbubble cavitation—stable cavitation and inertial cavitation. When insonated at low pressures, the microbubbles oscillate (ie, the microbubble repeatedly expands and contracts) in what is termed stable cavitation. – The resulting US waves scattered by the microbubbles occur not only at the applied frequency but also at harmonics of the applied frequency. When insonated at higher pressures, inertial cavitation begins to occur, where the microbubble oscillations become more pronounced and unstable, resulting in effects such as a lipid shedding and microbubble destruction. Above the pressure threshold for inertial cavitation, the waves emitted by a microbubble will occur across a broad band of frequencies. The nonlinear echo characteristics of stable and inertial cavitation allow the microbubble echoes to be separated from those of the surrounding tissue by selecting for harmonics. In the context of FUS, passive cavitation detection presents a unique opportunity to leverage the echoes generated by the FUS system to characterize cavitation. In this technique, the focus of a high frequency (typically >2 MHz) imaging US transducer is aligned with the focus of the FUS beam. This imaging transducer captures the echoes that are reflected during FUS application. Tissue echoes are separated from the oscillating microbubbles by using a receiving transducer that operates at a different frequency than the FUS transmitter or with signal processing techniques to filter the unique signal of the microbubbles. This technique can be extended into cavitation mapping by using an array of receivers that allow for beamforming to be performed to map and detect levels of cavitation across the field of view (see Fig. ). , Cavitation mapping has important applications in the field of BBB opening , and as a tool to noninvasively create qualitative maps of FUS acoustic fields. Feedback loops to control the level and type of cavitation are using these cavitation mapping techniques in preclinical systems. – These techniques are being evaluated for clinical neuronavigation-guided systems and may provide better control of BBB opening and histotripsy in the near future from companies including NaviFUS and Delsona. , Ultrasonic Thermometry A significant barrier to the adoption of FUS has been the need for high frame rate temperature mapping to guide thermal ablation. MR thermometry provides temperature data in HIFU systems that rely on temperature monitoring to determine when thermal ablation occurs. There have been many efforts to develop ultrasonic thermometry to measure the change in temperature of tissue during HIFU. The possible benefits of ultrasonic thermometry compared with MR thermometry include higher frame rates, lower cost, and convenient tracking and alignment with respect to the HIFU beam. Unfortunately, ultrasonic thermometry presents significant physical challenges including the nonlinear relationship between temperature change and speed of sound and the variation in speed of sound that occurs in fatty tissues. , Though an active field of research, it has not been used in the clinic due to these limitations.   Prior to the pairing of FUS with MRI performed by Hynynen in 1992, US was the dominant method for FUS guidance. , , In the context of FUS, US guidance presents the opportunity to image anatomy, image blood flow, image cavitation, measure a change in tissue properties, and monitor temperature changes. Though not all of these techniques are in clinical use, we will discuss each to provide insights into the benefits of US guidance and the direction of the field. Anatomical Imaging B-mode US guidance is used in the 3 indications most widely treated by FUS—uterine fibroids (36% of patients treated to date), liver tumors (21%), and prostate diseases (18%). By including a higher frequency imaging US transducer (>1 MHz) aligned with the low frequency FUS transducer (<1 MHz), a reliable detection of the FUS beam can be made in real time using pulse-echo imaging techniques. One model for imaging and treatment of the body is the table-based system, where the HIFU and imaging transducers are installed within a table with a water bath that is used to couple the transducers to the patient's body. With these platforms, used for uterine fibroids , and liver tumors, the treatment target can be selected using the B-mode US image as a guide. Then, as HIFU is applied, how well the tissue in the treatment site reflects the US waves (ie, echogenicity) can be monitored to determine if sufficient energy is being applied to cause ablation. A bright or hyperechoic spot, located at the focus of the HIFU beam, is the result of cavitation —the formation of gas microbubbles that then cause nonlinear reflections of the acoustic waves. In some cases, the B-mode image shows a hyperechoic bright spot at the site of tissue ablation after treatment, providing real-time imaging feedback of the ablation itself. , Although convenient, echogenicity alone is not a reliable method to determine the region of ablation or the extent of ablation. Prostate treatment uses a minimally invasive model, relying on a transrectal US probe that combines imaging and HIFU capability. The location of prostate cancer can be detected based on the hypoechoic (poorly reflected and appearing black on the US image) regions of interest, although this method of identifying tumors is highly operator dependent. , B-mode imaging is also used to guide histotripsy in the recently approved Histosonics Edison platform. Similar to HIFU systems used for thermal ablation, the Edison platform uses US guidance to plan and monitor treatment in real time. Following treatment with histotripsy, the tissue can exhibit a hypoechoic result (lacking US scattering), indicating liquefaction of the tissue. The detection of the cavitation cloud in real time is used to adjust the pressure amplitude of the applied US to induce the desired histotripsy effects (see Fig. ). , A similar approach is applied to select pressure amplitude in US-guided thermal ablation. Doppler Imaging—Blood Flow Doppler US imaging can be used to detect blood flow in real time. This offers the opportunity to monitor changes in tissue perfusion and activity that may allow detection of tumors as well as the effects of therapy on a tumor. – With FUS neuromodulation, Doppler US imaging has been proposed as a tool to monitor how the US modulates neuronal activity in real time. , Recent advances in high speed US imaging have opened up new opportunities to detect slow blood flow and higher-resolution imaging of blood flow—domains not previously able to be imaged with Doppler US.   B-mode US guidance is used in the 3 indications most widely treated by FUS—uterine fibroids (36% of patients treated to date), liver tumors (21%), and prostate diseases (18%). By including a higher frequency imaging US transducer (>1 MHz) aligned with the low frequency FUS transducer (<1 MHz), a reliable detection of the FUS beam can be made in real time using pulse-echo imaging techniques. One model for imaging and treatment of the body is the table-based system, where the HIFU and imaging transducers are installed within a table with a water bath that is used to couple the transducers to the patient's body. With these platforms, used for uterine fibroids , and liver tumors, the treatment target can be selected using the B-mode US image as a guide. Then, as HIFU is applied, how well the tissue in the treatment site reflects the US waves (ie, echogenicity) can be monitored to determine if sufficient energy is being applied to cause ablation. A bright or hyperechoic spot, located at the focus of the HIFU beam, is the result of cavitation —the formation of gas microbubbles that then cause nonlinear reflections of the acoustic waves. In some cases, the B-mode image shows a hyperechoic bright spot at the site of tissue ablation after treatment, providing real-time imaging feedback of the ablation itself. , Although convenient, echogenicity alone is not a reliable method to determine the region of ablation or the extent of ablation. Prostate treatment uses a minimally invasive model, relying on a transrectal US probe that combines imaging and HIFU capability. The location of prostate cancer can be detected based on the hypoechoic (poorly reflected and appearing black on the US image) regions of interest, although this method of identifying tumors is highly operator dependent. , B-mode imaging is also used to guide histotripsy in the recently approved Histosonics Edison platform. Similar to HIFU systems used for thermal ablation, the Edison platform uses US guidance to plan and monitor treatment in real time. Following treatment with histotripsy, the tissue can exhibit a hypoechoic result (lacking US scattering), indicating liquefaction of the tissue. The detection of the cavitation cloud in real time is used to adjust the pressure amplitude of the applied US to induce the desired histotripsy effects (see Fig. ). , A similar approach is applied to select pressure amplitude in US-guided thermal ablation.   Doppler US imaging can be used to detect blood flow in real time. This offers the opportunity to monitor changes in tissue perfusion and activity that may allow detection of tumors as well as the effects of therapy on a tumor. – With FUS neuromodulation, Doppler US imaging has been proposed as a tool to monitor how the US modulates neuronal activity in real time. , Recent advances in high speed US imaging have opened up new opportunities to detect slow blood flow and higher-resolution imaging of blood flow—domains not previously able to be imaged with Doppler US.   Contrast-enhanced US uses exogenous microbubbles to image the vasculature within the body. Due to their size, microbubbles can travel throughout the capillaries and other blood vessels of the body but do not pass into the interstitial spaces of the tissue. The nonlinear scattering of the microbubbles provides an opportunity to isolate their echo from that of the surrounding tissue to detect perfusion and blood vessel anatomy. The simplest techniques use the increased echogenicity where microbubbles are present to detect perfusion. Within the clinic, contrast-enhanced US has been used as an alternative to MRI to assess the nonperfused volume of uterine fibroids following ablation of the fibroid with HIFU. , Other opportunities for contrast-enhanced US techniques can be found in imaging other forms of cancer and the brain during BBB opening. Cavitation Detection Cavitation plays a significant role in multiple FUS treatments. To better characterize and control this mechanical bioeffect, different approaches have been adopted to selectively identify the cavitation activity of endogenous and exogenous microbubbles. This is possible due to the unique properties of the gas-filled microbubbles that scatter US waves. There are 2 domains of microbubble cavitation—stable cavitation and inertial cavitation. When insonated at low pressures, the microbubbles oscillate (ie, the microbubble repeatedly expands and contracts) in what is termed stable cavitation. – The resulting US waves scattered by the microbubbles occur not only at the applied frequency but also at harmonics of the applied frequency. When insonated at higher pressures, inertial cavitation begins to occur, where the microbubble oscillations become more pronounced and unstable, resulting in effects such as a lipid shedding and microbubble destruction. Above the pressure threshold for inertial cavitation, the waves emitted by a microbubble will occur across a broad band of frequencies. The nonlinear echo characteristics of stable and inertial cavitation allow the microbubble echoes to be separated from those of the surrounding tissue by selecting for harmonics. In the context of FUS, passive cavitation detection presents a unique opportunity to leverage the echoes generated by the FUS system to characterize cavitation. In this technique, the focus of a high frequency (typically >2 MHz) imaging US transducer is aligned with the focus of the FUS beam. This imaging transducer captures the echoes that are reflected during FUS application. Tissue echoes are separated from the oscillating microbubbles by using a receiving transducer that operates at a different frequency than the FUS transmitter or with signal processing techniques to filter the unique signal of the microbubbles. This technique can be extended into cavitation mapping by using an array of receivers that allow for beamforming to be performed to map and detect levels of cavitation across the field of view (see Fig. ). , Cavitation mapping has important applications in the field of BBB opening , and as a tool to noninvasively create qualitative maps of FUS acoustic fields. Feedback loops to control the level and type of cavitation are using these cavitation mapping techniques in preclinical systems. – These techniques are being evaluated for clinical neuronavigation-guided systems and may provide better control of BBB opening and histotripsy in the near future from companies including NaviFUS and Delsona. ,   Cavitation plays a significant role in multiple FUS treatments. To better characterize and control this mechanical bioeffect, different approaches have been adopted to selectively identify the cavitation activity of endogenous and exogenous microbubbles. This is possible due to the unique properties of the gas-filled microbubbles that scatter US waves. There are 2 domains of microbubble cavitation—stable cavitation and inertial cavitation. When insonated at low pressures, the microbubbles oscillate (ie, the microbubble repeatedly expands and contracts) in what is termed stable cavitation. – The resulting US waves scattered by the microbubbles occur not only at the applied frequency but also at harmonics of the applied frequency. When insonated at higher pressures, inertial cavitation begins to occur, where the microbubble oscillations become more pronounced and unstable, resulting in effects such as a lipid shedding and microbubble destruction. Above the pressure threshold for inertial cavitation, the waves emitted by a microbubble will occur across a broad band of frequencies. The nonlinear echo characteristics of stable and inertial cavitation allow the microbubble echoes to be separated from those of the surrounding tissue by selecting for harmonics. In the context of FUS, passive cavitation detection presents a unique opportunity to leverage the echoes generated by the FUS system to characterize cavitation. In this technique, the focus of a high frequency (typically >2 MHz) imaging US transducer is aligned with the focus of the FUS beam. This imaging transducer captures the echoes that are reflected during FUS application. Tissue echoes are separated from the oscillating microbubbles by using a receiving transducer that operates at a different frequency than the FUS transmitter or with signal processing techniques to filter the unique signal of the microbubbles. This technique can be extended into cavitation mapping by using an array of receivers that allow for beamforming to be performed to map and detect levels of cavitation across the field of view (see Fig. ). , Cavitation mapping has important applications in the field of BBB opening , and as a tool to noninvasively create qualitative maps of FUS acoustic fields. Feedback loops to control the level and type of cavitation are using these cavitation mapping techniques in preclinical systems. – These techniques are being evaluated for clinical neuronavigation-guided systems and may provide better control of BBB opening and histotripsy in the near future from companies including NaviFUS and Delsona. ,   A significant barrier to the adoption of FUS has been the need for high frame rate temperature mapping to guide thermal ablation. MR thermometry provides temperature data in HIFU systems that rely on temperature monitoring to determine when thermal ablation occurs. There have been many efforts to develop ultrasonic thermometry to measure the change in temperature of tissue during HIFU. The possible benefits of ultrasonic thermometry compared with MR thermometry include higher frame rates, lower cost, and convenient tracking and alignment with respect to the HIFU beam. Unfortunately, ultrasonic thermometry presents significant physical challenges including the nonlinear relationship between temperature change and speed of sound and the variation in speed of sound that occurs in fatty tissues. , Though an active field of research, it has not been used in the clinic due to these limitations.   Physiological processes may modify local heat conduction and energy absorption during heat deposition. Blood flow may increase during temperature increase and thus change heat conduction. In addition, the effect of heat stress on tissues is often delayed due to apoptotic processes and may thus be difficult to evaluate during or immediately following the intervention. Semiempirical relations between tissue damage and heat exposure have been derived by Sapareto and Dewey. To improve the therapeutic efficiency and the safety of the intervention, real-time mapping of temperature and thermal dose offer an excellent strategy to optimize such interventions and provide clinical therapy endpoints. , CT and US have been used to monitor thermal ablation. However, the extent of the lesion is difficult to assess during the intervention. Of the different imaging modalities, MRI appears to be the ideal tool for temperature mapping. Whereas temperature effects on MR measured parameters (resonance frequency, relaxation, diffusion) had been well described previously, the first report of temperature mapping by MRI appeared in 1983. The method was based on the longitudinal relaxation time of water protons (T 1 ). T 1 -based methods have been used in the clinic and preclinical applications, at low magnetic field. , Since then, alternative MR temperature imaging methods have been proposed such as techniques based on diffusion coefficient of water. However, the most widely used MR temperature mapping is based on temperature dependence of the water proton resonance frequency (PRF). – The excellent linearity of the temperature dependency of the PRF and its near-independence with respect to tissue type make the PRF-based approach the technique of choice, at mid to high field strength (see reviews – ). Quantification of the temperature changes allows endpoint determination based on thermal dose and temperature maps. Technical difficulties related to motion artifacts and quality control in MR temperature maps, and/or the need for real-time availability of temperature data have limited the use of temperature mapping. Improvements in MRI technologies have increased the accuracy and frequency of temperature mapping (see Fig. ). Promising new methods that limit motion-related artifacts have been developed. MR temperature mapping has already found widespread use in the clinic. The combination of FUS and MRI is especially powerful since MRI allows high precision target localization, and it is the only noninvasive imaging technique that has been shown to allow accurate and rapid temperature mapping in vivo at this time. As mentioned earlier, linearity and temperature sensitivity of the water PRF make PRF-based approach the preferred choice for FUS temperature monitoring. In the following section, principles of PRF-based MR thermometry are reviewed. Basic Principles of MR Temperature Mapping With the Water Proton Resonance Frequency Method The PRF methods usually employ rapid RF-spoiled gradient echo imaging. Following image processing, an MR signal is associated for each volume element as a complex number. Gray levels on anatomical images are proportional to the MR signal magnitude value, whereas its phase relates to the PRF. Temperature variation is calculated from phase changes for each voxel obtained at different times: Δϕ = γ · α · B 0 · ΔT · T E [1] where Δϕ is the phase difference, ΔT the temperature variation, γ the gyromagnetic ratio (~42.58.2π.10 6 rad/T), α the temperature coefficient (~0.01 ppm/K), T E the echo time, and B 0 the main magnetic field. The presence of lipids is a potential source of artifacts, since the temperature dependency of the resonance frequency of lipid hydrogen is near zero. In practice, lipid contribution to the MR signal can be conveniently suppressed in gradient echo imaging by frequency-selective slice excitation or alternative methods such as selective fat saturation or inversion recovery and multiecho methods combined with model-based reconstruction methods. High temperatures lead to coagulation and immediate cell death. However, moderate temperature elevations may also lead to cell death, albeit it with a delayed effect via biochemical pathways leading to apoptosis. Lethal effects of elevated temperatures have been studied by Sapareto and Dewey who established an empirical relation among temperature, duration of exposure, and cell death. Using such relations, temperature measurements throughout the procedure allow on-line thermal dose evaluation using the following equation: t eq = ∫ 0 t 2 T t − 43 dt if T > 43 ° C ∫ 0 t 4 T t − 43 dt if T < 43 ° C [2] Lethal dose is reached when t eq exceeds 240 minutes at 43°C. Thus, accurate thermal dose calculation requires precise thermometry since temperature has an exponential effect on thermal dose (see Equation 2). Requirements of MR Temperature for Real-Time Therapy Guidance and Endpoint Determination MR thermometry offers real-time therapy guidance for thermal therapies and therapy endpoints. However, in order to make MR temperature mapping widely useful for thermal therapies in clinical practice, much attention has to be paid on several conditions such as the following: (1) excellent spatial and temperature resolution; (2) on-line availability of accurate temperature maps and thermal dose maps; (3) robustness to artifacts in particular for temperature mapping of moving organs; (4) automatic quality assessment; and (5) on-line visualization of the hyperthermia procedure based on the presentation of thermal maps. High Temporal and Spatial Resolution A reliable temperature measurement must be performed on a region of interest defined by the radiologist. This region contains the ablation zone with a surrounding safety interval (typically at least 5 mm). Optimal MR thermometry requires high temporal and spatial resolution to precisely monitor temperature distribution within the targeted area and to predict the efficiency of the thermal treatment. Ideally, the time delay between 2 successive temperature maps should be small compared with the time in which the measured temperature change would have a significant influence on the therapy (a few seconds for ablation therapies for which temperatures above 50°C are reached, and up to several minutes at 43°C). Spatial resolution should be better than the smallest structure to be tracked individually (anatomically determined), such that temperature gradients across the voxel do not lead to signal intensity loss (due to phase distribution) and significant errors in the determination whether the lethal dose is reached (determined by the energy applicator and heat conduction). For FUS ablation, the resolution should be better than the dimensions of the focal point. Recent reviews have summarized developments in rapid imaging for temperature mapping. , Further acceleration of MRI has been demonstrated using spiral k-space coverage, simultaneous multislice sequences, reduction of k-space coverage, and prior knowledge of the region where temperature changes are expected. – Accurate Thermometry The accuracy of MR thermometry depends on several factors such as signal-to-noise ratio (SNR), echo time, field strength, and artifacts (see below). Acceleration of image acquisition to increase temporal resolution usually implies a SNR reduction. It is well established that electronic noise is Gaussian white noise. The MR system measures the real and imaginary components of the signal. As these 2 data each contain (ideally) Gaussian white noise, the resulting noise in the magnitude images will have a Rician distribution. , The precision of temperature estimation can thus be assessed from the standard deviation of a time series of measurements. Because the SNR increases with magnetic field intensity, magnetic field of 1 T (or higher) offers the possibility to obtain a thermal map with accuracy close to 1°C acquired with duration of 0.5 s per slice composed of 1 mm 3 voxels. Accurate thermometry allows the possibility of automatic control of the thermal procedure. Indeed, MR thermometry can be used to provide spatial and temporal temperature feedback control of the HIFU device. Low Level of Artifacts Due to Susceptibility Effects It has usually been assumed that, for PRF methods, macroscopic magnetic susceptibility effects are independent of temperature. A susceptibility field modification generates a perturbation of the magnetic field in the MRI and thus of the measured signal phase. Magnetic susceptibility is usually spatially inhomogeneous (eg, near intestines, stomach, esophagus, lungs) and increases with B 0 intensity. However, it has been demonstrated that temperature-dependent changes in magnetic susceptibility may not be negligible. , This is also the case when lipids are present. Methods have been presented for MRI thermometry in water and fat simultaneously. Low Level of Artifacts Due to Motion Complex organ displacement and deformation (mainly induced by breathing or cardiac activity) must be compensated in order to: - improve the precision of on-line temperature measurement by performing phase difference on pixels corresponding to identical spatial localization, - enable time-dependent information from images and thus make possible on-line thermal dose computation (as for each pixel, the history of the temperature corresponding to the same spatial localization in the organ is required, see Equation 2), and - preserve for each acquired image the validity of regions of interest defined by the radiologist at the beginning of the intervention (eg, target region and organs at risk). In addition, although the spatial transformation of the reference phase image can be corrected for upon tissue displacement, an unwanted phase shift may not be fully suppressed. In fact, the local magnetic susceptibility distribution may be modified. Therefore, both PRF-shift with temperature and tissue displacements may result in phase variations making it difficult to separate these 2 contributions within a single-phase image. Correction of motion-related errors is a challenging task in temperature MRI. Techniques are explored for correcting motion-related errors in PRF-based MR thermometry avoiding explicit modeling of the susceptibility field. Also, deep learning methods can be effectively employed to detect motion and correct for its effect on MRI temperature mapping. Accurate MR temperature monitoring has been demonstrated for immobilized organs such as the brain. Techniques have been developed to also allow MR temperature measurement of mobile organs. Much attention is paid to the organs of the abdomen as they are prone to periodic displacements induced by breathing or cardiac activity. Accelerated acquisition techniques are used (acquisition time per image faster than typical motion period), and the phase perturbation associated with organ motion is analyzed during a pretreatment step. A collection of multiple baseline reference phase images for different position of the organ is used to generate temperature maps. Alternatively, motion-related phase changes can be corrected for by using reference regions within each image. Visualization 3D visualization of temperature evolution based on multislice or volumetric imaging must be realized in order to allow a precise monitoring of the therapy. With recent improvements of spatial and temporal resolution of MR data, this visualization will become essential for the radiologist for an accurate estimation of ablation or hyperthermic areas and to avoid damage of healthy adjacent tissue.   The PRF methods usually employ rapid RF-spoiled gradient echo imaging. Following image processing, an MR signal is associated for each volume element as a complex number. Gray levels on anatomical images are proportional to the MR signal magnitude value, whereas its phase relates to the PRF. Temperature variation is calculated from phase changes for each voxel obtained at different times: Δϕ = γ · α · B 0 · ΔT · T E [1] where Δϕ is the phase difference, ΔT the temperature variation, γ the gyromagnetic ratio (~42.58.2π.10 6 rad/T), α the temperature coefficient (~0.01 ppm/K), T E the echo time, and B 0 the main magnetic field. The presence of lipids is a potential source of artifacts, since the temperature dependency of the resonance frequency of lipid hydrogen is near zero. In practice, lipid contribution to the MR signal can be conveniently suppressed in gradient echo imaging by frequency-selective slice excitation or alternative methods such as selective fat saturation or inversion recovery and multiecho methods combined with model-based reconstruction methods. High temperatures lead to coagulation and immediate cell death. However, moderate temperature elevations may also lead to cell death, albeit it with a delayed effect via biochemical pathways leading to apoptosis. Lethal effects of elevated temperatures have been studied by Sapareto and Dewey who established an empirical relation among temperature, duration of exposure, and cell death. Using such relations, temperature measurements throughout the procedure allow on-line thermal dose evaluation using the following equation: t eq = ∫ 0 t 2 T t − 43 dt if T > 43 ° C ∫ 0 t 4 T t − 43 dt if T < 43 ° C [2] Lethal dose is reached when t eq exceeds 240 minutes at 43°C. Thus, accurate thermal dose calculation requires precise thermometry since temperature has an exponential effect on thermal dose (see Equation 2).   MR thermometry offers real-time therapy guidance for thermal therapies and therapy endpoints. However, in order to make MR temperature mapping widely useful for thermal therapies in clinical practice, much attention has to be paid on several conditions such as the following: (1) excellent spatial and temperature resolution; (2) on-line availability of accurate temperature maps and thermal dose maps; (3) robustness to artifacts in particular for temperature mapping of moving organs; (4) automatic quality assessment; and (5) on-line visualization of the hyperthermia procedure based on the presentation of thermal maps. High Temporal and Spatial Resolution A reliable temperature measurement must be performed on a region of interest defined by the radiologist. This region contains the ablation zone with a surrounding safety interval (typically at least 5 mm). Optimal MR thermometry requires high temporal and spatial resolution to precisely monitor temperature distribution within the targeted area and to predict the efficiency of the thermal treatment. Ideally, the time delay between 2 successive temperature maps should be small compared with the time in which the measured temperature change would have a significant influence on the therapy (a few seconds for ablation therapies for which temperatures above 50°C are reached, and up to several minutes at 43°C). Spatial resolution should be better than the smallest structure to be tracked individually (anatomically determined), such that temperature gradients across the voxel do not lead to signal intensity loss (due to phase distribution) and significant errors in the determination whether the lethal dose is reached (determined by the energy applicator and heat conduction). For FUS ablation, the resolution should be better than the dimensions of the focal point. Recent reviews have summarized developments in rapid imaging for temperature mapping. , Further acceleration of MRI has been demonstrated using spiral k-space coverage, simultaneous multislice sequences, reduction of k-space coverage, and prior knowledge of the region where temperature changes are expected. – Accurate Thermometry The accuracy of MR thermometry depends on several factors such as signal-to-noise ratio (SNR), echo time, field strength, and artifacts (see below). Acceleration of image acquisition to increase temporal resolution usually implies a SNR reduction. It is well established that electronic noise is Gaussian white noise. The MR system measures the real and imaginary components of the signal. As these 2 data each contain (ideally) Gaussian white noise, the resulting noise in the magnitude images will have a Rician distribution. , The precision of temperature estimation can thus be assessed from the standard deviation of a time series of measurements. Because the SNR increases with magnetic field intensity, magnetic field of 1 T (or higher) offers the possibility to obtain a thermal map with accuracy close to 1°C acquired with duration of 0.5 s per slice composed of 1 mm 3 voxels. Accurate thermometry allows the possibility of automatic control of the thermal procedure. Indeed, MR thermometry can be used to provide spatial and temporal temperature feedback control of the HIFU device. Low Level of Artifacts Due to Susceptibility Effects It has usually been assumed that, for PRF methods, macroscopic magnetic susceptibility effects are independent of temperature. A susceptibility field modification generates a perturbation of the magnetic field in the MRI and thus of the measured signal phase. Magnetic susceptibility is usually spatially inhomogeneous (eg, near intestines, stomach, esophagus, lungs) and increases with B 0 intensity. However, it has been demonstrated that temperature-dependent changes in magnetic susceptibility may not be negligible. , This is also the case when lipids are present. Methods have been presented for MRI thermometry in water and fat simultaneously. Low Level of Artifacts Due to Motion Complex organ displacement and deformation (mainly induced by breathing or cardiac activity) must be compensated in order to: - improve the precision of on-line temperature measurement by performing phase difference on pixels corresponding to identical spatial localization, - enable time-dependent information from images and thus make possible on-line thermal dose computation (as for each pixel, the history of the temperature corresponding to the same spatial localization in the organ is required, see Equation 2), and - preserve for each acquired image the validity of regions of interest defined by the radiologist at the beginning of the intervention (eg, target region and organs at risk). In addition, although the spatial transformation of the reference phase image can be corrected for upon tissue displacement, an unwanted phase shift may not be fully suppressed. In fact, the local magnetic susceptibility distribution may be modified. Therefore, both PRF-shift with temperature and tissue displacements may result in phase variations making it difficult to separate these 2 contributions within a single-phase image. Correction of motion-related errors is a challenging task in temperature MRI. Techniques are explored for correcting motion-related errors in PRF-based MR thermometry avoiding explicit modeling of the susceptibility field. Also, deep learning methods can be effectively employed to detect motion and correct for its effect on MRI temperature mapping. Accurate MR temperature monitoring has been demonstrated for immobilized organs such as the brain. Techniques have been developed to also allow MR temperature measurement of mobile organs. Much attention is paid to the organs of the abdomen as they are prone to periodic displacements induced by breathing or cardiac activity. Accelerated acquisition techniques are used (acquisition time per image faster than typical motion period), and the phase perturbation associated with organ motion is analyzed during a pretreatment step. A collection of multiple baseline reference phase images for different position of the organ is used to generate temperature maps. Alternatively, motion-related phase changes can be corrected for by using reference regions within each image. Visualization 3D visualization of temperature evolution based on multislice or volumetric imaging must be realized in order to allow a precise monitoring of the therapy. With recent improvements of spatial and temporal resolution of MR data, this visualization will become essential for the radiologist for an accurate estimation of ablation or hyperthermic areas and to avoid damage of healthy adjacent tissue.   A reliable temperature measurement must be performed on a region of interest defined by the radiologist. This region contains the ablation zone with a surrounding safety interval (typically at least 5 mm). Optimal MR thermometry requires high temporal and spatial resolution to precisely monitor temperature distribution within the targeted area and to predict the efficiency of the thermal treatment. Ideally, the time delay between 2 successive temperature maps should be small compared with the time in which the measured temperature change would have a significant influence on the therapy (a few seconds for ablation therapies for which temperatures above 50°C are reached, and up to several minutes at 43°C). Spatial resolution should be better than the smallest structure to be tracked individually (anatomically determined), such that temperature gradients across the voxel do not lead to signal intensity loss (due to phase distribution) and significant errors in the determination whether the lethal dose is reached (determined by the energy applicator and heat conduction). For FUS ablation, the resolution should be better than the dimensions of the focal point. Recent reviews have summarized developments in rapid imaging for temperature mapping. , Further acceleration of MRI has been demonstrated using spiral k-space coverage, simultaneous multislice sequences, reduction of k-space coverage, and prior knowledge of the region where temperature changes are expected. –   The accuracy of MR thermometry depends on several factors such as signal-to-noise ratio (SNR), echo time, field strength, and artifacts (see below). Acceleration of image acquisition to increase temporal resolution usually implies a SNR reduction. It is well established that electronic noise is Gaussian white noise. The MR system measures the real and imaginary components of the signal. As these 2 data each contain (ideally) Gaussian white noise, the resulting noise in the magnitude images will have a Rician distribution. , The precision of temperature estimation can thus be assessed from the standard deviation of a time series of measurements. Because the SNR increases with magnetic field intensity, magnetic field of 1 T (or higher) offers the possibility to obtain a thermal map with accuracy close to 1°C acquired with duration of 0.5 s per slice composed of 1 mm 3 voxels. Accurate thermometry allows the possibility of automatic control of the thermal procedure. Indeed, MR thermometry can be used to provide spatial and temporal temperature feedback control of the HIFU device.   It has usually been assumed that, for PRF methods, macroscopic magnetic susceptibility effects are independent of temperature. A susceptibility field modification generates a perturbation of the magnetic field in the MRI and thus of the measured signal phase. Magnetic susceptibility is usually spatially inhomogeneous (eg, near intestines, stomach, esophagus, lungs) and increases with B 0 intensity. However, it has been demonstrated that temperature-dependent changes in magnetic susceptibility may not be negligible. , This is also the case when lipids are present. Methods have been presented for MRI thermometry in water and fat simultaneously.   Complex organ displacement and deformation (mainly induced by breathing or cardiac activity) must be compensated in order to: - improve the precision of on-line temperature measurement by performing phase difference on pixels corresponding to identical spatial localization, - enable time-dependent information from images and thus make possible on-line thermal dose computation (as for each pixel, the history of the temperature corresponding to the same spatial localization in the organ is required, see Equation 2), and - preserve for each acquired image the validity of regions of interest defined by the radiologist at the beginning of the intervention (eg, target region and organs at risk). In addition, although the spatial transformation of the reference phase image can be corrected for upon tissue displacement, an unwanted phase shift may not be fully suppressed. In fact, the local magnetic susceptibility distribution may be modified. Therefore, both PRF-shift with temperature and tissue displacements may result in phase variations making it difficult to separate these 2 contributions within a single-phase image. Correction of motion-related errors is a challenging task in temperature MRI. Techniques are explored for correcting motion-related errors in PRF-based MR thermometry avoiding explicit modeling of the susceptibility field. Also, deep learning methods can be effectively employed to detect motion and correct for its effect on MRI temperature mapping. Accurate MR temperature monitoring has been demonstrated for immobilized organs such as the brain. Techniques have been developed to also allow MR temperature measurement of mobile organs. Much attention is paid to the organs of the abdomen as they are prone to periodic displacements induced by breathing or cardiac activity. Accelerated acquisition techniques are used (acquisition time per image faster than typical motion period), and the phase perturbation associated with organ motion is analyzed during a pretreatment step. A collection of multiple baseline reference phase images for different position of the organ is used to generate temperature maps. Alternatively, motion-related phase changes can be corrected for by using reference regions within each image.   3D visualization of temperature evolution based on multislice or volumetric imaging must be realized in order to allow a precise monitoring of the therapy. With recent improvements of spatial and temporal resolution of MR data, this visualization will become essential for the radiologist for an accurate estimation of ablation or hyperthermic areas and to avoid damage of healthy adjacent tissue.   Transient formation of gas bodies during the action of HIFU within the biological tissues greatly improves energy deposition and efficacy of bioeffects. In this case, bubbles in the tissue are formed due to the action of US itself, during the passage of the rarefaction phase of the US pressure wave, from the air originally dissolved in blood and tissues, via inertial cavitation and rectified diffusion into the newly forming gas bodies. Newly formed gas bubbles rapidly compress and expand in the high-intensity US field, leading to localized energy deposition, both thermal and mechanical. When acoustic pressure is lower than the cavitation threshold, the passage of the US wave through the tissue occurs with minimal attenuation and energy loss, traditionally assumed as 0.3 dB/cm/MHz by FDA. Externally added “seed” gas bubbles are even more efficient at enhancing US energy deposition, as they do not require higher acoustic pressure values. Bioeffects of the combined action of medical US and administered gas-filled bubbles had been investigated initially in the context of safety of use of intravascular microbubble materials for contrast US imaging, about a quarter century ago. Skyba et al investigated US action on intravascular microbubbles in vivo, in a spinotrapezius muscle vasculature rat model. They varied energy transmitted by the US clinical imaging system transducer, within the range of peak negative acoustic pressures of 0.61 MPa and 1.53 MPa, also presented as Mechanical Index (MI) between 0.4 and 1.0, displayed on the system screen. MI is defined as peak negative acoustic pressure in MPa, divided by the square root of US frequency in MHz. It was found that, in response to the increase of acoustic pressure, more blood capillaries were broken, leading to higher level of RBC extravasation. We must note that the MI range investigated was within the approved MI limits for diagnostic imaging medical US systems in general. As the US contrast agent dose in rodents exceeded the human dose by 3 orders of magnitude, and only 0.015% of the microvessels were ruptured for each imaging frame, contrast agent safety was assured, and microbubble materials made their way into widespread clinical practice, with millions of injections administered to patients annually. The same group, Price et al also reported that the action of US on intravascular microbubbles (US+MB) in the rat muscle microscopy animal model can lead to extravascular delivery of not only RBCs but also submicron colloid particles, with the eventual goal of particle delivery outside of the vasculature, specifically in the insonated tissue. This publication opened up the field of US-assisted drug delivery across the endothelial barrier (see below). The next step in the assessment of the action of US on the microbubbles in the vasculature was made in the rat kidney model, by observation of petechial hemorrhages (ie, extravasation of RBCs) in response to US. It was found that at the highest acoustic pressures (MI ranged from 0.4 to 1.6) and lowest frequencies (1.8 to 6 MHz) the observed bioeffect (expressed by the blinded reader as the relative number of the observed kidney hemorrhages) was the highest for all the types of microbubbles tested. Unlike prior observations by Skyba et al and Price et al, in this study, intermittent US application (frame rate 1 Hz) was significantly more effective for the bioeffects generation, compared with continuous US (30 Hz). The latter observation could be attributed to microbubble hemodynamics in relation to the microbubble destruction rate: at continuous insonation, intravascular microbubbles could be destroyed in larger vessels, prior to entering the capillaries (Fig. A). At 1 Hz frame rate, the velocity of microbubbles within the blood flow is sufficiently fast to bring them from the larger vessels outside of the insonated area into the narrow capillaries. Expansion of microbubbles during the rarefaction phase of the US wave leads to the direct interaction with the narrow capillary vessel wall, with direct effect on the endothelial lining and possibly beyond that (Fig. B). Obviously, extravasation of RBCs is not a suitable scenario if US-assisted drug delivery is desired. Therefore, Hynynen et al and Sheikov et al investigated “softer” insonation scenarios, where microbubbles would be used as intravascular US cavitation enhancers, but the applied acoustic pressure would be significantly lower. The energy level in this case is not sufficient to break the vessel wall; however, a mild “massage” of the endothelial lining of the vessel wall by vibrating microbubbles is sufficient to achieve permeability enhancement. It is potentially highly important from the practical standpoint, specifically for the brain setting. Very few drugs are capable of crossing the BBB on their own. However, transient opening of the BBB that lasts from several minutes to several days became feasible, with the delivery of a wide array of drugs (including antibodies applied for tumor therapy and immunotherapy) as well as drug carriers (including drug-loaded liposomes, up to 100 nm in diameter). With proper adjustment of multielement transducer arrays, focusing of US may take place through the skull, with the focal zone as small as a grain of rice (see above, section that describes physics of US focusing). This allows completely noninvasive transcranial BBB opening specific to the desired area, for instance, brain tumor periphery, where the vasculature is sufficient to bring nutrients and drug with the blood flow, but the endothelium is still not sufficiently leaky to allow drug delivery across the BBB to the tumor. The only other requirement is that the drug carrier recirculates in the bloodstream for sufficient time and can recirculate through the vasculature with the open barrier, so it could extravasate and reach the brain. A clearly successful example of this approach was reported by Aryal et al. In their rodent glioma model after administration of doxorubicin-loaded long-circulating liposomes in combination with microbubbles and FUS, almost 40% of the animals were cured, that is, disease-free for 140 days. A very similar methodology, using an implanted US transducer, has recently also reached the clinical trial testing stage (see, eg, NCT05864534). The mechanisms of US+MB-induced BBB opening via changes exerted upon vascular endothelium depend on the dose of bubbles and acoustic parameters; they have been suggested very early by Sheikov et al and may include transcytosis, channel formation and fenestration, vesicle formation, vacuolization, and tight junction opening. Chen and Konofagou investigated extravasation of fluorescent dextrans with different molecular weights and found that acoustic pressure increase and change from stable cavitation to inertial cavitation of microvascular microbubbles influence the endothelial opening “pore” size. Caveolin-1 and clathrin-mediated endocytosis have been found to be involved in trans-BBB delivery. , Enhancement of Thrombolysis With Microbubbles and Ultrasound The use of US for the enhancement of clot dissolution was first described by Tachibana and Tachibana about 3 decades ago. It was found that a combination of microbubbles and US improved the activity of a thrombolytic agent. Since that time, over 200 articles have been published, and several clinical trials reported, so this field has moved closer to clinical practice. Therapy was achieved in several clot location scenarios, ranging from the peripheral vasculature to the coronary arteries and ischemic stroke. The general idea is that the vibration of microbubbles in the US field in close proximity or even within the thrombus may accelerate thrombolysis and/or reduce the required dose of a thrombolytic agent. The latter point is of particular importance, because current dosage of tissue plasminogen activator carries a small but strongly undesirable risk of bleeding, and a significant drug cost; both of these issues create limitations for the approach, and US seems to be able to address them. Furthermore, US treatment is reported to reduce microemboli in a thrombotic vascular occlusion model. All this may lead to a suggestion that US+MB treatment, if administered early, may be capable of reducing no-reflow scenario. However, care should be taken to avoid abrupt vasoconstriction by controlling insonation parameters. This approach has been investigated in a series of clinical trials, both in ischemic stroke (eg, NCT01678495) and myocardial infarction (NCT02410330) settings. These studies did not even employ complicated MR-guided FUS: a beneficial therapeutic bioeffect was established with the use of a clinical US imaging system, where higher acoustic pressure was applied under the guidance of diagnostic US imaging of myocardial perfusion defect. The clearly described example of the successful clinical testing is presented by Mathias et al, where statistically significant reduction of the size of myocardial infarction was observed for patients who received a combined US+MB treatment in addition to standard of care stenting. Targeted Transfection by Ultrasound Stimuli The general approach for nucleic acid delivery by a combination of microbubbles and US is (in hindsight) quite obvious: a microbubble is decorated with nucleic acid and administered in vivo, then US is applied in the area where transfection is desired. One of the explanations of this effect is that compression and expansion of bubbles on cell surface creates transient pores, so that nucleic acid can enter the cell cytoplasm. Very early on, this approach was presented by Porter et al who achieved a several-fold improvement of antisense oligonucleotide delivery by albumin-shell perfluorobutane microbubbles with oligonucleotide in an insonated canine kidney model, as compared with the noninsonated kidney. The simplest cationic bubble formulation consists of a lipid monolayer shell that contains a cationic lipid, for example, DSTAP (distearoyl trimethylammonium propane) and water-insoluble perfluorobutane gas. Nucleic acid is attached to the bubble surface by electrostatic forces. Following intravenous nucleic acid–bubble administration in vivo, with the immediate application of US to the area of desired delivery, efficient transfection following intravenous administration of the complexes is achieved. Transfection is most efficient in the tissues with high blood perfusion (eg, myocardium). At higher acoustic pressure values, nucleic acid delivery to the extravascular space is observed. At low acoustic pressure, for example, 100 kPa, transfection can be limited to endothelium only, without induction of BBB disruption or inflammation, see Figure . These DSTAP-plasmid electrostatically coupled microbubble formulations have also been shown to provide partial protection of nucleic acid from enzymatic degradation by nuclease. For the past 3 decades, this technique was the subject of significant preclinical investigations, with several hundred references listed in PubMed, but has not yet reached clinical testing. The main advantage of this approach is the selectivity of transfection. Numerous gene therapy and nucleic acid delivery approaches that apply systemic administration most often lead to efficient transfection in the liver, which is not always the intended and desired target organ. Yet the combination of nucleic acid and high melting point nonfusogenic lipid-based cationic bubbles only allow intracellular delivery within the insonated spot, with no transfection observed in any other organs, such as the liver or lungs (see Fig. ). As an alternative to the bubbles that carry nucleic acid directly, a clinical grade microbubble US contrast agent can be co-administered with nucleic acid (plasmid, small interfering RNA or microRNA). Nucleic acid can be complexed with a carrier protectant, for example, liposomes, LNP, or cationic polymer. But even a combination of existing clinical grade materials has not been brought to the clinical testing stage at this time. Ultrasound-Assisted Drug Delivery and Therapy Enhancement There are 2 general approaches that could be considered for the use of US in drug delivery. First is a co-administration, or subsequent administration, of microbubbles, followed by US action, and intravascular administration of the drug. Most of the study designs of this kind are rapidly progressing and have reached clinical trial stage, some demonstrating more successful therapeutic interventions than others. – This may be a “low-hanging fruit” because clinical-grade US contrast agent microbubbles are already approved by FDA, as are the drugs (most likely anticancer drug formulations), so the approval process that is required to reach a clinical trial stage is relatively straightforward, and associated cost is not as high. The potential mechanisms for this approach have been discussed earlier in this review, where US+MB action on the vasculature is discussed; general principles are presented in a recent review by Durham et al. The second approach suggests the preparation of sonosensitive drug carrier particles. This alternative is yet to reach clinical testing, because the US-triggered particle is a novel chemical entity, microbubbles or microbubble complexes that carry the drug. It is either the microbubbles that carry hydrophobic pharmaceuticals, such as rapamycin , in their lipid shell, or a hydrophilic drug, incorporated inside liposomes that are placed onto the microbubble shell. This approach may imply a more complicated formulation and costly approval process. Long-term stability of the research-grade formulations is also not always assured. However, some preclinical successes have been achieved in vitro , and in vivo. There is another issue of the currently available gas-filled microbubble formulations: limited circulation time. Typical clinical gas-filled microbubbles, even with the poorly soluble perfluorocarbon gas core, circulate in the bloodstream at the desired concentrations for no longer than a minute or two. Microbubbles deflate, and gas is removed by exhalation via lungs. Thus, the locally triggered drug release and localized delivery procedure is limited by the amount of the drug carrier that passes through the insonated target region of interest in several minutes. Still, if the drug critical toxicity is avoided by the drug construct, and therapeutic index, that is, the ratio of toxicity level (eg, LD50, median lethal dose) to efficacy (eg, EC50, median effective dose) is improved, it may still be beneficial to use this technique as compared with the free drug. Free drugs in cancer chemotherapy may clear from the bloodstream as quickly, but are known to cause critical toxicity, for example, in bone marrow and cardiac muscle. Longer circulation time may still be feasible for some other sonosensitive formulations. One of the early designs was associated with the triggering of superheated nanoparticles (nanodroplets) made of low-boiling liquid fluorocarbons with US. Early studies with US contrast perfluoropentane nanodroplets (Echogen) that dated back to 1990s were reviewed in detail by Lin and Pitt. These early generation nanodroplets were stabilized with a Zonyl fluorosurfactant, phased out due to environmental and safety concerns, and were not approved by FDA. More recent studies that provided detailed experimental evidence and understanding of the observed phenomena were performed with the nanodroplets stabilized with biocompatible and safe phospholipid shells, and the core of liquid perfluoropropane, perfluorobutane, and perfluoropentane. Despite the low boiling points of these bulk perfluorocarbons, and their high vapor pressure, when placed in a nanoparticle core, these compounds stay liquid at physiological temperature, until conversion to gas takes place by the application of a US pulse of sufficient energy. The nanodroplet sensitivity to vaporization by US is better with the lowering of the perfluorocarbon boiling point and increase of the nanodroplet particle diameter. It was suggested that fluorocarbon boiling is induced by superharmonic focusing of the US wave, with the newly created boiling point located within the droplet. Nanodroplets stabilized with a lipid monolayer shell may stay in the circulation for longer time than microbubbles, for example, about 3-fold longer for decafluorobutane. Another approach, which uses the capability of FUS to achieve mild tissue hyperthermia, is based on thermosensitive liposomes, for example, those known as Thermodox, developed by Celsion. To achieve release upon mild hyperthermia, dipalmitoyl phosphatidylcholine with a phase transition temperature of 41°C was originally selected as the liposome membrane lipid. To create Thermodox, this design was subsequently improved, by adding PEG (polyethylene glycol)–lipid and monoacyl lipid, to improve circulation time and achieve faster content efflux at the elevated temperature. These particles release entrapped doxorubicin only when they pass through the tissue, heated by US to 41°C. The circulation time of Thermodox liposomes is shorter when compared with the original Doxil/Lipodox long-circulating PEG-decorated liposomes, but much longer than for the free drug. It took almost 2 decades to move this formulation from early preclinical research to clinical trials. , It is quite possible that hybrid systems, where nanodroplets are combined with another drug carrier, will be able to achieve extended circulation time. For instance, drug-loaded liposomes that co-entrap nanodroplets and drug demonstrate US-triggered contents release. , One other approach is the use of “hitchhiking”: nanoparticles are known to stay in the bloodstream for a longer period than their “normal” circulation time, if they are adhered to red blood cells prior to injection in the bloodstream. Therefore, we may expect that perfluorocarbon nanodroplet-decorated RBCs will stay in the bloodstream, while it is already shown that a single microsecond-long FUS pulse leads to the explosive conversion of an RBC-bound nanodroplet, RBC membrane poration and drug content release within a fraction of a second (see Fig. ). Drugs incorporated into the sonosensitive particles do not have to be the small molecules as described above. Proteins, such as antibodies or enzymes, can be sequestered as well, that is, separated from the biological milieu by the microcapsule shell (eg, a liposome bilayer membrane). The entrapped protein is unable to act upon the biological target, until US is applied, and liposome contents released, so that enzymatic action on the biologically relevant substrate could take place. One such example is a US-induced thrombin release from microbubble-liposome complexes, which could be eventually tested for externally triggered blood vessel clotting. In an “inverted” version of the antitumor therapy contents release approach, an inactive prodrug was administered systemically first. Then local high-pressure FUS treatment destroyed the cells in the tumor target area, leading to the localized release of intracellular enzymes, which convert the nontoxic prodrug, doxorubicin-glucuronide, to the therapeutically active free doxorubicin. Induction of Vascular Inflammation by Insonation of Microbubbles When microbubbles in the vasculature were subjected to a US field in the course of BBB opening experiments, sterile inflammatory response with acute inflammation and immune activation was observed. The dose and type of microbubbles as well as the acoustic parameters of FUS could be of importance. Overall, it may be possible and desirable to achieve BBB opening without the activation of the proinflammatory NFkB pathway. This may be useful, for example, when BBB opening treatment of Alzheimer’s disease is the intended goal. However, tight control of the dose of US contrast agent, type of the agent (eg, lipid microbubbles vs albumin-shell bubbles, having very different size distribution), and even the type of carrier gas used for inhalation (oxygen vs medical air) may all influence the outcome. , One may also note that avoiding inflammation is not always desirable. On the contrary, we may be able to take advantage of inducing localized vascular endothelium activation in the target cancer tissue upon insonation in order to increase accumulation of the innate immunity cells in the insonated tumor, which might help convert “cold” tumors to “hot” ones, with a desirable antitumor immune response. Tumor Vasculature Shutdown: Combination of Ultrasound and Microbubbles An observation was made by Bunte et al that treatment of a subcutaneous tumor mass in a murine model with a physical therapy US system led to the reduction of tumor blood flow. It was later traced to the presence of US contrast microbubbles in the vasculature. Similar observations were reported by Hwang et al. Combined MB+US treatment could result in a transient (minutes) blood flow stoppage in the tumor vasculature, but repeated MB+US treatments resulted in tumor suppression without additional chemotherapy. Todorova et al describe prolonged antivascular effects in the insonated tumors, especially if combined with chemotherapy. A specific term “angiotripsy” has been suggested. The mechanism of blood flow cessation was in some cases traced to platelet activation in the insonated vasculature. One might hypothesize that enhanced clotting in tumor vasculature might be related to a high level of prothrombotic phosphatidylserine on the surface of vascular endothelium in the tumors. To enhance therapeutic efficacy, this approach may be combined with other types of localized therapeutic intervention, such as radiotherapy. A combination of US antivascular therapy and radiotherapy has already reached the clinical trial stage, reported, for example, by Eisenbrey et al and a recently completed NCT04431674.   The use of US for the enhancement of clot dissolution was first described by Tachibana and Tachibana about 3 decades ago. It was found that a combination of microbubbles and US improved the activity of a thrombolytic agent. Since that time, over 200 articles have been published, and several clinical trials reported, so this field has moved closer to clinical practice. Therapy was achieved in several clot location scenarios, ranging from the peripheral vasculature to the coronary arteries and ischemic stroke. The general idea is that the vibration of microbubbles in the US field in close proximity or even within the thrombus may accelerate thrombolysis and/or reduce the required dose of a thrombolytic agent. The latter point is of particular importance, because current dosage of tissue plasminogen activator carries a small but strongly undesirable risk of bleeding, and a significant drug cost; both of these issues create limitations for the approach, and US seems to be able to address them. Furthermore, US treatment is reported to reduce microemboli in a thrombotic vascular occlusion model. All this may lead to a suggestion that US+MB treatment, if administered early, may be capable of reducing no-reflow scenario. However, care should be taken to avoid abrupt vasoconstriction by controlling insonation parameters. This approach has been investigated in a series of clinical trials, both in ischemic stroke (eg, NCT01678495) and myocardial infarction (NCT02410330) settings. These studies did not even employ complicated MR-guided FUS: a beneficial therapeutic bioeffect was established with the use of a clinical US imaging system, where higher acoustic pressure was applied under the guidance of diagnostic US imaging of myocardial perfusion defect. The clearly described example of the successful clinical testing is presented by Mathias et al, where statistically significant reduction of the size of myocardial infarction was observed for patients who received a combined US+MB treatment in addition to standard of care stenting.   The general approach for nucleic acid delivery by a combination of microbubbles and US is (in hindsight) quite obvious: a microbubble is decorated with nucleic acid and administered in vivo, then US is applied in the area where transfection is desired. One of the explanations of this effect is that compression and expansion of bubbles on cell surface creates transient pores, so that nucleic acid can enter the cell cytoplasm. Very early on, this approach was presented by Porter et al who achieved a several-fold improvement of antisense oligonucleotide delivery by albumin-shell perfluorobutane microbubbles with oligonucleotide in an insonated canine kidney model, as compared with the noninsonated kidney. The simplest cationic bubble formulation consists of a lipid monolayer shell that contains a cationic lipid, for example, DSTAP (distearoyl trimethylammonium propane) and water-insoluble perfluorobutane gas. Nucleic acid is attached to the bubble surface by electrostatic forces. Following intravenous nucleic acid–bubble administration in vivo, with the immediate application of US to the area of desired delivery, efficient transfection following intravenous administration of the complexes is achieved. Transfection is most efficient in the tissues with high blood perfusion (eg, myocardium). At higher acoustic pressure values, nucleic acid delivery to the extravascular space is observed. At low acoustic pressure, for example, 100 kPa, transfection can be limited to endothelium only, without induction of BBB disruption or inflammation, see Figure . These DSTAP-plasmid electrostatically coupled microbubble formulations have also been shown to provide partial protection of nucleic acid from enzymatic degradation by nuclease. For the past 3 decades, this technique was the subject of significant preclinical investigations, with several hundred references listed in PubMed, but has not yet reached clinical testing. The main advantage of this approach is the selectivity of transfection. Numerous gene therapy and nucleic acid delivery approaches that apply systemic administration most often lead to efficient transfection in the liver, which is not always the intended and desired target organ. Yet the combination of nucleic acid and high melting point nonfusogenic lipid-based cationic bubbles only allow intracellular delivery within the insonated spot, with no transfection observed in any other organs, such as the liver or lungs (see Fig. ). As an alternative to the bubbles that carry nucleic acid directly, a clinical grade microbubble US contrast agent can be co-administered with nucleic acid (plasmid, small interfering RNA or microRNA). Nucleic acid can be complexed with a carrier protectant, for example, liposomes, LNP, or cationic polymer. But even a combination of existing clinical grade materials has not been brought to the clinical testing stage at this time.   There are 2 general approaches that could be considered for the use of US in drug delivery. First is a co-administration, or subsequent administration, of microbubbles, followed by US action, and intravascular administration of the drug. Most of the study designs of this kind are rapidly progressing and have reached clinical trial stage, some demonstrating more successful therapeutic interventions than others. – This may be a “low-hanging fruit” because clinical-grade US contrast agent microbubbles are already approved by FDA, as are the drugs (most likely anticancer drug formulations), so the approval process that is required to reach a clinical trial stage is relatively straightforward, and associated cost is not as high. The potential mechanisms for this approach have been discussed earlier in this review, where US+MB action on the vasculature is discussed; general principles are presented in a recent review by Durham et al. The second approach suggests the preparation of sonosensitive drug carrier particles. This alternative is yet to reach clinical testing, because the US-triggered particle is a novel chemical entity, microbubbles or microbubble complexes that carry the drug. It is either the microbubbles that carry hydrophobic pharmaceuticals, such as rapamycin , in their lipid shell, or a hydrophilic drug, incorporated inside liposomes that are placed onto the microbubble shell. This approach may imply a more complicated formulation and costly approval process. Long-term stability of the research-grade formulations is also not always assured. However, some preclinical successes have been achieved in vitro , and in vivo. There is another issue of the currently available gas-filled microbubble formulations: limited circulation time. Typical clinical gas-filled microbubbles, even with the poorly soluble perfluorocarbon gas core, circulate in the bloodstream at the desired concentrations for no longer than a minute or two. Microbubbles deflate, and gas is removed by exhalation via lungs. Thus, the locally triggered drug release and localized delivery procedure is limited by the amount of the drug carrier that passes through the insonated target region of interest in several minutes. Still, if the drug critical toxicity is avoided by the drug construct, and therapeutic index, that is, the ratio of toxicity level (eg, LD50, median lethal dose) to efficacy (eg, EC50, median effective dose) is improved, it may still be beneficial to use this technique as compared with the free drug. Free drugs in cancer chemotherapy may clear from the bloodstream as quickly, but are known to cause critical toxicity, for example, in bone marrow and cardiac muscle. Longer circulation time may still be feasible for some other sonosensitive formulations. One of the early designs was associated with the triggering of superheated nanoparticles (nanodroplets) made of low-boiling liquid fluorocarbons with US. Early studies with US contrast perfluoropentane nanodroplets (Echogen) that dated back to 1990s were reviewed in detail by Lin and Pitt. These early generation nanodroplets were stabilized with a Zonyl fluorosurfactant, phased out due to environmental and safety concerns, and were not approved by FDA. More recent studies that provided detailed experimental evidence and understanding of the observed phenomena were performed with the nanodroplets stabilized with biocompatible and safe phospholipid shells, and the core of liquid perfluoropropane, perfluorobutane, and perfluoropentane. Despite the low boiling points of these bulk perfluorocarbons, and their high vapor pressure, when placed in a nanoparticle core, these compounds stay liquid at physiological temperature, until conversion to gas takes place by the application of a US pulse of sufficient energy. The nanodroplet sensitivity to vaporization by US is better with the lowering of the perfluorocarbon boiling point and increase of the nanodroplet particle diameter. It was suggested that fluorocarbon boiling is induced by superharmonic focusing of the US wave, with the newly created boiling point located within the droplet. Nanodroplets stabilized with a lipid monolayer shell may stay in the circulation for longer time than microbubbles, for example, about 3-fold longer for decafluorobutane. Another approach, which uses the capability of FUS to achieve mild tissue hyperthermia, is based on thermosensitive liposomes, for example, those known as Thermodox, developed by Celsion. To achieve release upon mild hyperthermia, dipalmitoyl phosphatidylcholine with a phase transition temperature of 41°C was originally selected as the liposome membrane lipid. To create Thermodox, this design was subsequently improved, by adding PEG (polyethylene glycol)–lipid and monoacyl lipid, to improve circulation time and achieve faster content efflux at the elevated temperature. These particles release entrapped doxorubicin only when they pass through the tissue, heated by US to 41°C. The circulation time of Thermodox liposomes is shorter when compared with the original Doxil/Lipodox long-circulating PEG-decorated liposomes, but much longer than for the free drug. It took almost 2 decades to move this formulation from early preclinical research to clinical trials. , It is quite possible that hybrid systems, where nanodroplets are combined with another drug carrier, will be able to achieve extended circulation time. For instance, drug-loaded liposomes that co-entrap nanodroplets and drug demonstrate US-triggered contents release. , One other approach is the use of “hitchhiking”: nanoparticles are known to stay in the bloodstream for a longer period than their “normal” circulation time, if they are adhered to red blood cells prior to injection in the bloodstream. Therefore, we may expect that perfluorocarbon nanodroplet-decorated RBCs will stay in the bloodstream, while it is already shown that a single microsecond-long FUS pulse leads to the explosive conversion of an RBC-bound nanodroplet, RBC membrane poration and drug content release within a fraction of a second (see Fig. ). Drugs incorporated into the sonosensitive particles do not have to be the small molecules as described above. Proteins, such as antibodies or enzymes, can be sequestered as well, that is, separated from the biological milieu by the microcapsule shell (eg, a liposome bilayer membrane). The entrapped protein is unable to act upon the biological target, until US is applied, and liposome contents released, so that enzymatic action on the biologically relevant substrate could take place. One such example is a US-induced thrombin release from microbubble-liposome complexes, which could be eventually tested for externally triggered blood vessel clotting. In an “inverted” version of the antitumor therapy contents release approach, an inactive prodrug was administered systemically first. Then local high-pressure FUS treatment destroyed the cells in the tumor target area, leading to the localized release of intracellular enzymes, which convert the nontoxic prodrug, doxorubicin-glucuronide, to the therapeutically active free doxorubicin.   When microbubbles in the vasculature were subjected to a US field in the course of BBB opening experiments, sterile inflammatory response with acute inflammation and immune activation was observed. The dose and type of microbubbles as well as the acoustic parameters of FUS could be of importance. Overall, it may be possible and desirable to achieve BBB opening without the activation of the proinflammatory NFkB pathway. This may be useful, for example, when BBB opening treatment of Alzheimer’s disease is the intended goal. However, tight control of the dose of US contrast agent, type of the agent (eg, lipid microbubbles vs albumin-shell bubbles, having very different size distribution), and even the type of carrier gas used for inhalation (oxygen vs medical air) may all influence the outcome. , One may also note that avoiding inflammation is not always desirable. On the contrary, we may be able to take advantage of inducing localized vascular endothelium activation in the target cancer tissue upon insonation in order to increase accumulation of the innate immunity cells in the insonated tumor, which might help convert “cold” tumors to “hot” ones, with a desirable antitumor immune response.   An observation was made by Bunte et al that treatment of a subcutaneous tumor mass in a murine model with a physical therapy US system led to the reduction of tumor blood flow. It was later traced to the presence of US contrast microbubbles in the vasculature. Similar observations were reported by Hwang et al. Combined MB+US treatment could result in a transient (minutes) blood flow stoppage in the tumor vasculature, but repeated MB+US treatments resulted in tumor suppression without additional chemotherapy. Todorova et al describe prolonged antivascular effects in the insonated tumors, especially if combined with chemotherapy. A specific term “angiotripsy” has been suggested. The mechanism of blood flow cessation was in some cases traced to platelet activation in the insonated vasculature. One might hypothesize that enhanced clotting in tumor vasculature might be related to a high level of prothrombotic phosphatidylserine on the surface of vascular endothelium in the tumors. To enhance therapeutic efficacy, this approach may be combined with other types of localized therapeutic intervention, such as radiotherapy. A combination of US antivascular therapy and radiotherapy has already reached the clinical trial stage, reported, for example, by Eisenbrey et al and a recently completed NCT04431674.   Focusing an externally applied acoustic field to the desired areas of the body for therapy is a rapidly expanding area of research. In some instances, it has moved from the early stage applied physics research and development to a clinical trial setting, and even approved clinical indications, such as noninvasive essential tremor treatment. Most successes and clinical translation currently take place in the image-guided hyperthermia domain, especially under MRI guidance. Real-time temperature monitoring appears to be currently the best strategy for sustained improvement of the outcome of FUS procedures. For optimal treatment, temperature in the targeted volume should be controlled accurately and continuously. With the expanded development of nonthermal mechanisms of action, additional imaging techniques will be important to the continued success of FUS. In this regard, US imaging of cavitation and blood flow may provide important tools for new mechanisms of action that rely on cavitation and nonthermal effects. In addition, B-mode US guidance provides a low-cost alternative to MRI thermometry that monitors cavitation and echogenicity to select treatment parameters and identify changes in the tissue due to the action of FUS. Some areas of research, such as triggered drug and gene delivery, are progressing slower, due to enhanced complexity and cost of generating robust formulations that would bring success in clinical trials. Overall, the progress already achieved and expected to occur in the area of FUS therapy is greatly enhanced by real-time imaging guidance, provided by medical imaging.   