General Principles of the MRgFUS Method

By clinical medicine standards, the MRgFUS method can be considered relatively “young,” but the preliminary outcomes of its 15-year development and remarkable achievements in treating nervous system diseases are quite impressive. Before presenting the main capabilities, results, practical experience, and prospects of using MRgFUS in neurology, it seems appropriate to provide a brief analysis of the general principles of using ultrasound for therapeutic purposes.

History of Surgery with Focused Ultrasound

Ultrasound refers to mechanical vibrations and waves in elastic media in the frequency range (20,000 Hz – 10 ¹⁰ Hz). Low-frequency ultrasound (2×10 ⁴ – 10⁵ Гц), средних (10 ⁵ – 10⁷ Гц) и высоких частот (10 ⁷ – 10¹⁰ Hz). In medicine, it is customary to distinguish low-intensity ultrasound (0.125–3.0 W/cm ²) and high intensity (from 3.0 W/cm ²) (Reznikov et al., 2015).

Еще в конце XVI century Italian Lazzaro Spallanzani suggested that bats use inaudible sounds for orientation in the dark, but in practice, the use of ultrasound began, as is often the case, with military purposes: in 1916, submarines started using sonar to detect enemy ships (Reznikov et al., 2015). Also, in the first half of the 20th century, the medical application of ultrasonic waves began. The first mentions of ultrasound in medicine appear in 1918, and of the therapeutic significance of ultrasonic waves in 1939, when they began to be used for the treatment of inflammatory diseases of muscles and joints Patzold, 1918; Pohlman et al., 1939).

The use of focused ultrasound waves for intracerebral ablation was first described J.G. Lynn et al. (1942). Later F.J. Fry (1958) developed a complex device for its time with 4 piezoelectric transducers for the safe ablation of intracranial tumors by performing a craniotomy to create a “window” for transmitting acoustic waves. Since then, for more than half a century, similar attempts to use ultrasonic radiation as the “ideal tool” for brain impact have been made repeatedly, however, despite some successes ( Zubiani, 1951; Lindstrom, 1954; Fry, Meyers, 1962; Hynynen et al., 1993), until the late 1990s, the main problem had not been solved—the passage of ultrasound through the skull bones. All operations were performed after trepanation of the skull without opening the dura mater, whereas the passage of focused ultrasound through the skull bones caused a significant reduction in ultrasound strength and heating of the bone. This problem was solved using multiple synchronized sources of ultrasonic radiation, which were evenly placed on a hemisphere around the head; none of them caused significant heating of the skull and brain tissues along the beam path (Tyurnikov, Gushcha, 2016). Additionally, each source had a controller that adjusted the phase shift so that the ultrasonic waves from all sources reached the target point simultaneously and acted briefly Hynynen, Jolesz, 1998; Clement et al., 2000). As a result, the heating occurred with such a high gradient that just 3 mm from the target point, the effect of the ultrasound was negated, and the tissue temperature did not increase. To reduce the heating of the skin and skull bones, the space between the helmet with radiation sources and the patient’s head, limited by an elastic membrane, was filled with circulating cold water.

Other technical improvements have been proposed (Tyurnikov, Gushcha, 2016). For instance, while the initial systems of focused ultrasound used ultrasound imaging for treatment, which has limited targeting accuracy and does not allow real-time temperature measurement at the target, the combination of focused ultrasound with MRI has significantly improved the control of the ablation focus and enabled precise thermometry to assess the degree of impact Hardy et al., 1994).

The procedure parameters began to be adjusted (in the intervals between focused ultrasound pulses) based on the obtained MRI image. Temperature monitoring in the destruction zone provides the method with maximum efficiency and safety ( Hynynen et al., 1993; Ricke, Butts Pauly, 2008).

The result of many years of work by numerous researchers was the system ExAblate 4000 (InSightec), consisting of a helmet with 1024 piezoelectric elements, which can be used to exert thermal effects on the brain with very high precision by focusing ultrasound waves at a frequency of 650 kHz (Fig. 2.1). It is integrated into a single complex with an MRI scanner of 1.5 or 3 Tesla and allows optimal direction of acoustic energy to a strictly localized target, as well as control of heating temperature and thermal dose at the target point. Thus, the precision of ablation with high-intensity focused ultrasound is ensured firstly by intraoperative management through MRI scanning and real-time thermometry, and secondly by using a phased array antenna to control the ultrasound beam.

Pilot therapeutic procedures using MRgFUS in 2009 were described in patients with neuropathic pain, but the first approved indication for clinical use of MRgFUS in 2016 was ET ( Martin et al., 2009; Osterholt et al., 2020; Abe et al., 2021). In 2018, the use of the method for the tremor form of Parkinson’s disease was approved ( Martínez-Fernández et al., 2020; Perlmutter, Ushe, 2020; Gallay et al., 2021; Lin et al., 2021), and later for a whole range of other diseases. To this day, movement disorders remain the leading area of application for MRgFUS in neurology.

Table 2.1 presents the main historical milestones in the development of MRgFUS technology and its application for the treatment of movement disorders.

Table 2.1 The history of the development of MRgFUS technology and its application for the treatment of movement disorders

Mechanisms of Focused Ultrasound Impact on the Body

Focused ultrasound, in the most general sense, is a technology that allows the concentration of ultrasound effects at a specific point in a particular organ. This enables the induction of various effects in tissues, which has therapeutic potential for many clinical conditions ( Foley et al., 2013). The control of the focus area when using this technology is carried out in real-time using a special thermometry program during MRI (Galimova et al., 2020)

Physical factors determining the impact of focused ultrasound on tissues can be divided into thermal and mechanical. Typically, in practice, they are combined with each other, although they can have isolated effects. The result of the physical impact of ultrasound is always a certain biological effect, which depends not only on the power, duration, and mode (continuous or pulsed) of exposure but also on the specific tissue (target). For example, muscle and bone react differently to ultrasound due to their different acoustic parameters.

Thermal exposure Continuous delivery of acoustic energy increases the temperature at the focal point of the waves inside the human body ( Al-Bataineh et al., 2012). The increase in temperature is due to the absorption of energy from ultrasound waves in a viscous medium, as well as the scattering of ultrasound as it passes through mechanical inhomogeneities in the tissue. The magnitude and duration of the temperature increase are referred to as the “thermal dose” delivered to the tissue. Ultrasound energy can be used to create a low level of hyperthermia lasting from minutes to hours (local hyperthermia) ( Jang et al., 2010). On the contrary, short and localized exposure to high temperatures destroys tissues by denaturing proteins and forming coagulative necrosis ( Webb et al., 2011).

As a rule, to achieve the effect of thermal coagulation of brain tissue, an ultrasound energy intensity of about 1000 W/cm² is used (Kholyavin, 2019). Since part of the ultrasound energy is lost during its passage to the target point and calculating the amount of heat delivered to the intracerebral target is usually impossible, the main parameter used to regulate the size of the destruction focus is the temperature in the exposure zone. It has been established that raising the temperature at a certain point in the brain to 57°C leads to complete cell death within 1 second, and at a temperature of 54°C C – within 3 seconds Schlesinger et al., 2017). Ultrasound heating of tissue above 65–70° C not recommended, as in this case there may be dilation of capillaries with the formation of less dense necrosis and an increased risk of tissue hemorrhages (Kholyavin, 2019)

Mechanical impact. High-intensity ultrasound pulses lead to significant pressure changes in the target tissue, which, in turn, cause numerous effects—from vibration to cavitation ( Jang et al., 2010; Uchida et al., 2012).

The most significant mechanical effect of focused ultrasound is cavitation – the oscillation of gas bubbles as ultrasound waves pass through tissues ( Uchida et al., 2012). These dissolved gases can be either introduced microbubbles (for example, to open the blood-brain barrier (BBB) – see chapter 9) or microbubbles created by the ultrasound itself due to negative pressure ( Chen et al., 2009; Deng, 2010). If the transmitted energy is sufficiently high, the rapid and intense pulsation of microbubbles leads to their collapse (non-stationary or collapsing cavitation), the emergence of areas with a sharp local increase in pressure, temperature rise, and tissue destruction. Low-energy ultrasound leads to stable cavitation, causing transient damage to cell membranes and increasing their permeability.

The cavitation process is less controllable compared to the thermal effect of ultrasound exposure (Kholyavin, 2019). Reducing the frequency of ultrasound lowers the energy intensity threshold at which cavitation occurs. Thus, by increasing the frequency of ultrasound exposure, the risk of cavitation can be reduced or avoided. The introduction of microbubbles (used as a contrast agent in ultrasound imaging) allows for more controllable, stable cavitation at relatively low ultrasound power.

The constant time-averaged force acting on an object in an acoustic field is called the acoustic radiation force. This force arises from the transfer of motion from the sound field to the object ( Rooney, Nyborg, 1972). Using ultrasound, it is possible to create tissue microflows in biological objects ( Karazapryanov, Titianova, 2016). Radial forces acting on cells in the field of ultrasound waves are more complex: they are associated with the “standing wave” effect and allow particles to be collected in lines. This is used to determine, for example, the erythrocyte sedimentation rate or to displace contrast agents to the walls of blood vessels in laboratory animals (Gurbatov, Klemina, 2011; Dayton et al., 1999).

The power of acoustic radiation is the primary physical mechanism leading to phenomena such as radial torque, acoustic streaming, acoustic levitation, and acoustic fountain effects. Radiation torque causes the rotation of symmetrical particles, while asymmetrical particles may predominantly rotate in an orientation corresponding to the acoustic field. An acoustic field propagating in a fluid medium can induce a bulk flow of liquid, known as acoustic streaming. The flow rate depends on several factors—absorption coefficient, speed of sound, kinematic viscosity, intensity, beam geometry, and nonlinear propagation ( Dalecki, 2004). The induction of fluid flow in an acoustic field, caused by the impact of ultrasonic waves, can trigger apoptosis processes responsible for the formation of delayed cell destruction. Additionally, the mechanical impact of ultrasound on cell membranes affects the functioning of ion channels and thus participates in the proposed neuromodulatory effect of transcranial ultrasound (Kholyavin, 2019)

By combining biological effects, various results can be achieved. The most commonly used is the thermal effect of tissue destruction and protein denaturation ( Webb et al., 2011; Al-Bataineh et al., 2012). Increase in temperature to a relatively low (42° C) and maintaining it for several minutes can increase blood supply and drug absorption in the target region without irreversible tissue damage ( Wang et al., 2010; Thanou, Gedroyc, 2013). Mechanical effects are used to destroy regions (histotripsy), as well as to deliver drugs to the target and create pores with ultrasound (sonoporation) ( Liang et al., 2010). Stable cavitation creates an acoustic stream that increases fluid flow in the cellular environment. This acceleration of flow can help open pores and direct drug molecules into the cell Wu et al., 2002; Collis et al., 2010). Focused ultrasound can reversibly increase the permeability of the vascular wall (especially when interacting with administered microbubbles), allowing chemicals to pass through it and penetrate the surrounding tissue Deng, 2010). A similar effect was used for the temporary opening of the BBB and the delivery of various compounds to the brain ( Tung et al., 2011; Park et al., 2012).

Thus, the impact of focused ultrasound on body tissues through interconnected thermal and mechanical factors leads to a variety of biological effects, such as thermocoagulation, histotripsy, modulation of cell functions, thrombolysis, sonoporation, increased permeability of histohematological barriers, etc., which are used in various methods of treatment and diagnosis of human diseases.

Modern Capabilities and Features of MRgFUS Impact on the Brain

As mentioned above, in neurology and neurosurgery, the use of focused ultrasound is particularly challenging due to the need to pass through the skull bones, which absorb and reflect a significant amount of ultrasonic energy, disrupting the focusing process. To overcome this effect, it has been proposed to use transducers with a large number of high-energy sources. Circulation of cool water around the patient’s head helps prevent thermal damage to the skin, and the hemispherical design of the helmet itself allows for more transmitters and proper distribution of thermal energy.

Another issue with transcranial passage of ultrasound beams is their significant aberration. The uneven thickness of the skull and the inner plate of the cranial bones, combined with the high speed of sound waves within bone tissue, leads to the defocusing of ultrasound beams ( Hynynen et al., 2006). To solve this problem, a computerized multichannel phased array sensor in the form of a hemisphere is currently used. The direction of each beam from the transducer is controlled by computer calculations and adjusted for varying skull thickness based on CT scanning. In combination with acoustic modeling, this allows phase adjustment to focus the waves on a small area ( Hynynen et al., 2006).

If we summarize the main technological innovations that contributed to the practical implementation of the MRgFUS method for brain surgeries, we can highlight 5 fundamental solutions implemented in the setup ExAblate (Kholyavin, 2019):

1) the presence of a large number of ultrasound sources (1024 sources) evenly distributed around the patient’s skull, directing ultrasound waves convergently into the target area of the brain
2) guiding the ultrasound beam to target points in the brain using an MRI scanner integrated with a device generating ultrasound energy
3) preliminary CT scan of the patient’s brain to create a virtual spatial model of the skull and verify the thickness of its bones, as well as phase-amplitude correction for each ultrasound source (using large phased-array ultrasound transducers and adaptive focusing methods). This allows effective compensation for the defocusing of ultrasound energy caused by waves passing through the skull bones;
4) conducting MRI thermometry during the therapeutic session to obtain temperature maps of the patient’s brain, ensuring real-time temperature control at the target site. The temperature after processing MRI images is measured by the shift in proton resonance frequency using fast gradient echo sequences, and the exposure volume is determined using T2-weighted fast spin echo sequences ( Hynynen et al., 1996);
5) conducting test exposures (test sonications) at the target point in the brain with heating in the temperature range of reversible tissue damage (41–45° C) under the control of the patient’s neurological status. Short-term heating to the specified temperatures leads to the functional deactivation of the target area without causing tissue necrosis

Thanks to points 1–3, precise targeting of the intracerebral target is ensured, and points 4–5 allow for feedback between the impact and the treatment result. This makes it possible to adjust the focus of the impact during the treatment session if necessary, achieving the maximum therapeutic effect without adverse side effects. When the optimal position of the target in the desired zone is established, confirmed by the effect of trial (reversible) sonications, decisive heating to a temperature of 51–64° is performed C, what leads to the formation of a therapeutic ablation focus. One ultrasound pulse lasts several seconds, and the controlled destruction of brain tissue is achieved with high precision (the average error is no more than 0.50–0.75 mm). Considering the various sizes of target structures in different situations of functional stereotaxis, a therapeutic session can result in a destruction focus with a diameter of 2–3 to 10 mm.

The ability to perform trial, reversible effects on the brain using ultrasound at sub-therapeutic intensity levels while simultaneously assessing the patient’s response is a significant advantage of the MRgFUS method compared to standard stereotactic radiofrequency ablation or the use of a gamma knife. Trial sonication, which does not cause irreversible tissue damage, allows for the desired clinical effect (e.g., suppression of tremor) to be observed or warns of possible complications (paresis, dysarthria). This ensures high safety of the method and creates conditions for achieving better results. Moreover, adjusting the position of the target in the brain does not increase the invasiveness of the operation, as is observed in invasive stereotactic interventions, where changing the target position requires inserting an electrode into the brain along a new trajectory.

In studies on biological models, a clear demarcation of the coagulation necrosis focus from the surrounding tissue was identified, with no tissue damage outside the impact zone (Kholyavin, 2019 Cohen et al., 2007). On the MRI of the brain of patients in mode T2-weighted images immediately after treatment in the target area show 3 concentric zones: a hypointense zone I in the center of the lesion, surrounded by a hyperintense zone II, which, in turn, is surrounded peripherally by a slightly hyperintense zone III. Zones I и II represent necrotic tissue and areas of cytotoxic edema, whereas the zone III reflects the perifocal edema of the surrounding intact brain tissue. At the boundary of the zones II и III a thin hyperintense rim can be observed, caused by the accumulation of hemosiderin. Zone III disappears within 1–7 days, whereas the zones I и II evolve into a cyst of round or oval shape within 1–4 weeks. It has been established that the diameter of the area II and the size of the destruction in the target zone located in the gray matter of the brain, 1 day after exposure, corresponds to the size of the 51° isotherm contour C according to MRI thermometry data conducted during the treatment session (Kholyavin, 2019; Bond, Elias, 2018). For this condition, the exposure should be at a temperature not lower than 51° C must last more than 3 seconds. For targets located in the area of the brain’s conductive pathways (white matter), which are more resistant to thermal damage, exposure at the specified temperatures must be repeated 4 times. This is explained by the dense arrangement of axons and their protection by the myelin sheath ( Magara et al., 2014).

Thus, the main advantage of MRgFUS technology compared to other types of functional impacts on the brain is obvious – the absence of the need for a craniotomy. When treating with MRgFUS, there is no risk of infection or hemorrhagic complications, no implants in the brain tissue, and no need for programming and constant monitoring of the stimulator’s condition (as is the case with DBS). This procedure does not require a sterile operating room or anesthesia (which allows for constant contact with the patient during the treatment session and necessary adjustments to the focus for maximum therapeutic effect) and is essentially performed on an outpatient basis. A clear result is immediately visible and achieved in one session. It is important to note that most possible side effects (usually mild) occur directly during the procedure and generally resolve on their own within a few days (Galimova et al., 2020; Gallay et al., 2018). All the features of MRgFUS significantly reduce the duration of patients’ stay in medical facilities and ease the postoperative recovery period overall.

It is no coincidence that the analysis of the economic component revealed a statistically significantly higher efficiency of using MRgFUS technology compared to DBS and stereotactic radiosurgery ( Ravikumar et al., 2017). The projected costs for therapy using MRgFUS were significantly lower compared to the aforementioned methods, despite the high cost of the equipment and consumables used (membrane and gel for sealing the fluid cooling the patient’s head, screws for its fixation, etc.), as well as the fact that the procedure occupies the MRI scanner for 3–6 hours. Summarizing the experience gained in various clinics, it can be concluded that the economic advantages of the MRgFUS method are achieved due to the absence of:

• prolonged hospitalization
• anesthetic management
• severe complications
• implantation of devices into the body and their subsequent replacement;
• the need for repeated visits to adjust the generator parameters

Alongside the undeniable advantages of MRgFUS, like any other method, it has its limitations and drawbacks (Tyrnikov, Gusha, 2016). In particular, the procedure cannot be performed on patients with thick skull bones or well-developed spongy bone tissue of the skull, previously installed intracerebral implants, foreign bodies in the brain, etc. Additionally, the system has limitations in generating high temperatures at points more than 5 cm away from the intercommissural line. In the work W.S. Chang et al. (2015) for these reasons, during thalamotomy using MRgFUS, 28% of patients could not complete the operation. A certain disadvantage of MRgFUS is the inability to conduct MER, as well as micro and macrostimulation during the operation – according to some authors, macrostimulation is considered more controllable and informative compared to trial sonication ( Palur et al., 2002).

Thanks to all the above-mentioned advantages, the MRgFUS technology, having made significant progress over 15 years, has secured a strong position in the treatment of patients with chronic progressive CNS diseases that are indications for stereotactic interventions on the brain. It is one of the proven highly effective methods for treating movement disorders (ET, PD, dystonia, Holmes tremor), successfully used in patients with neuropathic pain of central and peripheral origin. Clinical experience is gradually accumulating in the treatment of obsessive-compulsive disorder, depression, epilepsy, and some neuro-oncological syndromes using MRgFUS. Early-stage clinical research is underway in the application of MRgFUS for temporary opening of the BBB (targeted delivery of chemotherapeutic agents to tumor tissue, introduction of stem cells or genetically engineered constructs into target brain areas, increasing the availability of antibodies to pathological proteins in neurodegenerative diseases), sonothrombolysis in patients with ischemic stroke, neuromodulation, treatment of hydrocephalus, and more. (Kholyavin, 2019; Mainprize et al., 2016; Baek et al., 2022; Samuel et al., 2023). The most significant of these developments are discussed in detail in the following chapters.