Keywords

MRI, Focused ultrasound, MRgFUS, BBB, Neurological disorders, Drug delivery

Introduction

The BBB serves as a selective barrier between the bloodstream and the central nervous system (CNS), composed of tightly joined endothelial cells and efflux transporters. This barrier permits the passage of ions and small lipid-soluble molecules under 400 Da while restricting larger molecules [1,2]. The BBB is crucial for maintaining a stable CNS environment, safeguarding it from potentially harmful substances, but it also poses a significant challenge for delivering therapeutic agents to the brain [3,4]. Various strategies to enhance BBB permeability are being explored, broadly classified into transcellular and paracellular methods [5]. Transcellular techniques involve increasing the lipophilicity of molecules or utilizing carrier-mediated transport to facilitate BBB crossing [6]. However, these methods are constrained by the pharmacological properties of the agents. Paracellular strategies involve the temporary disruption of tight junctions, using either chemical agents like hyperosmolar compounds or physical methods such as ultrasound, which can transiently open the BBB [5].

Among the paracellular approaches, the use of low-intensity ultrasound combined with microbubble injection has gained prominence. This method allows for targeted, reversible BBB opening in specific brain regions, demonstrated in various preclinical models [7]. Clinical trials utilizing MRgFUS, also known as MR-HIFUS, have demonstrated promising outcomes, showcasing their efficacy and safety in permeabilizing the BBB. This technique employs an MRI-compatible helmet with focused ultrasound transducers, providing real-time imaging guidance and allowing for the monitoring of BBB disruption via gadolinium (Gd) contrast [1]. MRgFUS is particularly beneficial in treating movement disorders and certain psychiatric conditions, with potential applications in neuromodulation when used at lower intensities [8-11]. Additionally, devices like SonoCloud offer a more continuous method of BBB disruption, suitable for treatments requiring repeated sessions [12,13].

The blood-tumor barrier (BTB) presents an additional challenge and opportunity for targeted drug delivery. As brain tumors grow, they disrupt the normal architecture of the BBB, forming the BTB, which is characterized by increased permeability [14,15]. However, this disruption is often insufficient for therapeutic concentrations of drugs to reach the tumor site without affecting the surrounding healthy brain tissue. MRgFUS offers a promising method for enhancing drug delivery across the BTB by transiently increasing its permeability, thereby improving the localization of chemotherapeutic agents and minimizing systemic toxicity [14,16].

Following BBB and BTB opening, systemic administration of drugs can achieve targeted delivery to the CNS, including small molecules, monoclonal antibodies, and even neural stem cells [17]. This approach is under investigation for various conditions, including Alzheimer's disease, Parkinson's disease, and brain tumors [18,19]. The efficacy of MRgFUS in enhancing the delivery of chemotherapeutic agents has been demonstrated, showing significantly higher concentrations in the brain tissue with reduced systemic toxicity [16]. This review focuses on the latest advancements in ultrasound-mediated BBB and BTB permeabilization techniques, emphasizing their potential for safe, targeted, and effective drug delivery in the treatment of Neurological disorders.

Mechanisms and Techniques in MRgFUS

High-intensity focused ultrasound

High-intensity focused ultrasound (HIFUS) is an innovative noninvasive method that offers an alternative to traditional hyperthermia techniques. It works by concentrating ultrasonic waves from an external transducer to raise the temperature in specific tissues, providing targeted thermal effects. This technique is particularly beneficial as it can be precisely directed to deep tissues within the body, achieving both mild and ablative hyperthermia. The interaction between ultrasound waves and tissue can produce thermal and mechanical bioeffects, enhancing drug extravasation, uptake, and release from temperature-sensitive or pressure-sensitive drug delivery systems [20-22].

These bioeffects can also improve intracellular drug delivery and increase tissue permeability [23]. The parameters of HIFUS, such as power, frequency, duty cycle, and timing, can be adjusted to fine-tune and control these bioeffects. When combined with MRI for guidance, known as MRgFUS or MR-HIFUS, this technique allows for high-resolution anatomical imaging and precise monitoring of temperature changes and tissue displacement during therapy.

MR-HIFUS has been extensively used for the ablation of symptomatic uterine leiomyomata [24,25]. Its application is expanding into various oncological treatments, including benign breast fibroadenomas [26], malignant breast carcinoma [27], prostate cancer [28], palliative care for painful bone metastases [29], and brain tumor ablation [30]. Additionally, MRgFUS/MR-HIFUS is being explored for treating neurological disorders such as epilepsy, essential tremor, neuropathic pain, and Parkinson's disease [31-33]. It also shows potential as an immunomodulatory tool to prevent local tumor recurrence and metastasis post-ablation [34]. Moreover, MRgFUS/MR-HIFUS has demonstrated effectiveness in enhancing drug delivery to the brain by modulating the BBB [35].

Principles of focused ultrasound

Ultrasound is a type of mechanical wave that travels through a medium at frequencies higher than the human audible range of 20 Hz [36]. When ultrasound waves encounter a boundary between two media with different acoustic impedances, phenomena such as reflection and refraction occur [37]. Specular reflection, which happens when ultrasound waves hit a smooth surface, helps delineate interfaces between soft tissues, thereby providing clear images of organ boundaries for diagnostic purposes [38,39]. Additionally, ultrasound waves experience attenuation due to absorption and scattering, which can convert acoustic energy into heat. This thermal effect can lead to protein denaturation, coagulation, cell necrosis, and ultimately, tissue ablation [30,40,41]. Ultrasound can also induce cavitation, a process where microbubbles form and oscillate, potentially causing mechanical damage to tissues through high-pressure and shear forces [42,43].

The development of FUS technology began with Gruetzmacher's design of a curved quartz plate to focus ultrasound beams at a specific point in 1935 [44]. Since then, advances in phased array transducers, comprising hundreds of piezoelectric elements, have significantly improved FUS's ability to navigate complex tissue structures, such as bone, and to focus on multiple points simultaneously, enhancing the therapeutic volume [45,46]. Modern systems allow clinicians to adjust sonication parameters, including acoustic energy intensity, frequency, and duration, tailored to the treatment's objectives and patient characteristics.

The thermal and mechanical effects of ultrasound underpin its therapeutic applications in clinical settings, particularly in stereotactic ablative surgery. This technique, known as MRgFUS, offers a non-invasive alternative to traditional thermal therapies like radiofrequency (RF) ablation. Furthermore, under specific sonication conditions, cavitation effects can selectively and reversibly open the BBB, which is being actively investigated for its potential in enhancing drug delivery to the brain and neuromodulation [10,47,48].

MRgFUS BBB opening procedure

The MRgFUS BBB opening procedure utilizes the 220 kHz ExAblate Neuro 4000 system, featuring 1024 ultrasound transducers integrated with a 3T MRI scanner. Initially, the subject's hair is shaved closely to the scalp, and a Cosman-Roberts-Wells stereotactic frame is applied under local anesthesia for precise alignment. The subject is positioned supine on the MRI table, equipped with compression stockings, and all pressure points are carefully padded. The frame is then secured to the ExAblate helmet, with degassed water acting as a medium between the scalp and transducers to ensure optimal contact. Vital signs and electrocardiogram are continuously monitored using MR-compatible, non-invasive systems throughout the procedure. Participants remain comfortable and alert, with the ability to halt the procedure if they experience discomfort or an emergency [36,49].

The procedure commences with the acquisition of planning MR images, registered with baseline scans for accurate targeting. The focus is on opening the BBB in specific regions of the body, such as the arm or leg. Two targets are identified in these regions, each carefully chosen to avoid critical areas and minimize the risk of adverse effects. Within these targets, the focused ultrasound beam is precisely guided to ensure effective and safe BBB opening, following a predetermined pattern to maximize coverage while minimizing risk.

For transcranial focused ultrasound procedures targeting the BBB, estimating in vivo tissue pressure for individual patients is complicated by variations in skull shape, density, and the specific locations targeted. To address this, the optimal power level for BBB opening is established using cavitation feedback through a process known as a ramp test. This technique involves gradually increasing the power in small increments, usually around 5%, during brief sonications. The device's hydrophones monitor for sub-harmonic acoustic signals from the target area, which indicate the onset of cavitation. Once detected, the optimal power for opening the BBB is determined to be approximately 50% of this cavitation threshold, ensuring effective treatment while reducing the risk of side effects.

The procedure includes administering one to two 90-second ultrasound cycles to achieve the desired effect, accompanied by the intravenous injection of microbubbles. The dose of microbubbles is carefully controlled to not exceed a specific limit. Ultrasound is applied in burst mode with specific parameters, including a pulse repetition period and duty cycle, targeting multiple areas.

During the procedure, real-time monitoring includes acoustic monitoring, MRI thermometry, and direct feedback from the patient. MRI thermometry is performed using a standard fast spoiled gradient echo (FSPGR) sequence, ensuring no interference with the cavitation receivers in the focused ultrasound array. After each ultrasound cycle, patients undergo clinical and radiological assessments to check for adverse events, using Gd-enhanced T1-weighted and gradient echo (GRE) sequences to evaluate BBB permeability, tissue integrity, and the presence of microhemorrhages. The appearance of new Gd enhancement in the targeted brain area indicates successful BBB opening, at which point the procedure is concluded. Clinical and MRI assessments are conducted the following day to evaluate the reversibility of BBB opening and any adverse events [49].

Mechanism of BBB disruption using ultrasound

Microbubbles, which are small gas-filled spheres with a microsphere shell, are introduced into the bloodstream peripherally. Initially designed as a contrast agent for cardiac ultrasound diagnostics, microbubbles are now used for opening the BBB via ultrasound [1,50]. When microbubbles are exposed to ultrasound in targeted brain regions, as precisely defined by patient-specific MRI, they oscillate, causing a temporary disruption of the BBB (Figure 1). The degree of BBB disruption depends on factors such as acoustic pressure, duration of sonication, and microbubble size [51,52]. Excessive oscillation forces can, however, lead to hemorrhage. These forces induce stretching, acoustic streaming, and shear stress on blood vessels, impacting the permeability of tight junctions and the activity of efflux transporter proteins [48,53].

There are various microbubble formulations available, with the ideal type promoting stable cavitation, reducing P-glycoprotein production nearby, and facilitating caveolae formation (membrane invaginations) [54,55]. A study evaluating three different commercial microbubble products found them equally effective in terms of the extent and duration of BBB permeability [56]. The performance of microbubbles can be optimized by adjusting ultrasound parameters such as power and duration based on the microbubble's size and half-life [56].

Early preclinical research was crucial in establishing safe parameters for FUS. These studies indicated that using lower frequencies and peak pressures is consistently safe, promoting harmonic bubble oscillations instead of bubble collapse, thereby preventing damage to blood vessels, neurons, or glial cells [57]. A mechanical index (MI), calculated as the peak negative pressure (estimated in situ ) divided by the square root of the ultrasound frequency, was developed, with an MI ≤ 0.45 being consistently safe and not associated with hemorrhage [58].

The process of BBB permeability is dynamic, occurring almost immediately after the ultrasound sonication of microbubbles. Confirmation of BBB closure can be achieved via histology and non-invasive imaging, such as MRI, typically within 3-24 hours [35,59]. T1-weighted MRI with gadolinium contrast is commonly used to demonstrate BBB opening and subsequent restoration. Initial human trials focused on a small tissue volume (1 cm ³ ) per treatment, but recent studies have safely increased this to 40 cm ³ with repeated sessions [60], with even larger volumes being tested in non-human primates [61].

Alternative strategies using ultrasound to enhance drug delivery are also under active investigation. One such approach involves nanodroplets, which have a longer half-life than microbubbles and can carry therapeutic agents. These nanodroplets can be administered systemically and, upon reaching the ultrasound target area, vaporize to release the drug locally [62]. This technique has been used, for instance, to deliver phenobarbital to the amygdala for treating agitation in an Alzheimer's disease mouse model [63]. Another innovative method involves administering a piezoelectric nanogenerator to rodents, which embeds into neuronal membranes. When activated by FUS, this device stimulates tyrosine hydroxylase activity, enhancing dopamine production in striatal neurons [64]. While nanodroplets and nanogenerators hold great potential for future innovations, the established safety and reliability of transcranial and implantable FUS devices for BBB opening have already led to several completed and ongoing translational human clinical trials [1] (Figure 1).

Advanced Techniques in Image-Guided Drug Delivery

In the context of personalized medicine, achieving the "right treatment for the right patient at the right time" is a fundamental principle. To achieve this, innovative tools are necessary to tailor therapies to individual patient needs. Image-guided drug delivery (IGDD) has emerged as a promising approach, utilizing clinical imaging techniques to enhance the precision of drug delivery systems. This method involves the use of imaging to delineate target and non-target areas, as well as for screening, treatment planning, monitoring, and post-treatment evaluation [65].

MR-guided drug delivery

MR imaging is particularly advantageous in IGDD due to its capacity for high spatial and temporal resolution imaging and quantitative assessments during therapeutic interventions. MR-guided drug delivery not only complements existing minimally invasive therapies but also holds the potential to improve their efficacy and broaden their clinical applications. This review delves into the current advancements in MR-guided drug delivery, with an emphasis on hyperthermia-mediated delivery techniques and prospective future developments in this field [65].

MR imaging offers distinct advantages over other imaging modalities like ultrasonography and computed tomography, making it particularly suitable for guiding drug delivery. One key advantage is its ability to provide excellent contrast between soft tissues and between normal and abnormal structures, owing to tissue-specific MR parameters. This high level of contrast enhances the precision of treatment planning. Importantly, MR imaging does not involve ionizing radiation, which is crucial for procedures requiring repeated imaging to monitor drug delivery and assess tumor progression [65].

Additionally, MR imaging offers comprehensive anatomical, functional, and metabolic data through volumetric and multiplanar imaging techniques. This multifaceted capability has been increasingly employed in the planning, monitoring, and post-treatment assessment of drug delivery systems, particularly when used in conjunction with HIFUS [66].

Clinical Applications of MRgFUS in Neurological Disorders

Brain tumor management with MRgFUS

The utilization of MRgFUS and microbubbles has shown promise in overcoming the BBB and BTB for the enhanced delivery of chemotherapeutics. In a rat brain glioma model, MRgFUS significantly increased the concentration of doxorubicin (DOX) in gliomas, achieving levels over 2.5 times higher than control tumors at one-hour post-treatment, and nearly 14 times higher at 24 hours [14,67]. This effect suggests that MRgFUS not only facilitates drug delivery but also prolongs the duration of drug concentration in the target area. However, the response of the BBB and BTB to MRgFUS can vary depending on the systemic treatment used. For instance, while MRgFUS improved BCNU delivery to normal brain tissue by 340%, the increase was only 202% in tumor-bearing rats (p < 0.05) [68], indicating a possible treatment-dependent variation in BTB disruption. These findings are supported by unpublished experimental data in tumor-implanted mice (Figure 2).

The application of MRgFUS can significantly enhance the penetration of chemotherapeutic drugs through the BTB, thereby improving treatment efficacy and survival outcomes in tumor models. For instance, in a rat glioma model, combining MRgFUS with systemic carmustine (BCNU) administration resulted in a notable increase in median survival time compared to untreated controls (53 days vs . 29 days) and those receiving BCNU alone (53 days vs . 32 days, with no reported significance) [68]. This suggests that MRgFUS enhances the effectiveness of chemotherapy, as it did not provide survival benefits on its own. In a similar study, the combination of temozolomide (TMZ) and MRgFUS in rats significantly extended median survival compared to no treatment (23 days vs . 20 days), whereas TMZ alone did not show a significant difference compared to the untreated group (21 days vs . 20 days) [69]. Additionally, the use of MRgFUS with TMZ significantly reduced tumor growth compared to TMZ alone, with tumors growing 21 times their original size without MRgFUS and only 5 times with MRgFUS over a one-week period. These findings indicate that MRgFUS not only improves the delivery of systemic chemotherapy but also enhances its overall efficacy, although further research is needed to fully understand the dose-response relationship.

Nanoparticles provide an alternative to liposomal encapsulation, which may cause potential side effects [70]. Magnetic nanoparticles offer a promising synergy with MRgFUS for drug delivery. These magnetic nanoparticles can be linked to chemotherapeutics and directed to specific areas using an external magnetic field, a technique known as magnetic targeting [71]. Additionally, magnetic nanoparticles serve as contrast agents, potentially eliminating the need for gadolinium in MRgFUS treatments [72]. While magnetic nanoparticles typically struggle to cross the BBB, their accumulation in brain tissue can be significantly enhanced when combined with MRgFUS [73]. This process involves an "open-then-pull" method, where MRgFUS first disrupts the BBB, allowing circulating MNPs to be attracted to the permeabilized area by a magnetic field.

Early animal studies have demonstrated the potential of combining magnetic nanoparticles with MRgFUS to enhance treatment delivery to brain tumors. For example, in a rat glioma model treated with epirubicin-conjugated magnetic nanoparticles, MRgFUS, and magnetic targeting, there was a 2.4-fold increase in magnetic nanoparticle accumulation in the targeted hemisphere compared to the opposite hemisphere, whereas the combination of magnetic nanoparticles and MRgFUS without magnetic targeting resulted in only a 1.2-fold increase [73]. This enhanced accumulation led to a 16-fold increase in epirubicin concentration in brain tissue treated with MRgFUS and MT compared to MRgFUS alone. Additionally, the group receiving magnetic nanoparticles MRgFUS, and magnetic targeting showed prolonged survival compared to untreated controls (31 days vs . 18 days) or those treated with magnetic nanoparticles and MRgFUS without magnetic targeting (30.5 days vs . 20 days). Tumor growth was also significantly delayed in the MRgFUS and MT group compared to untreated controls over seven days (106% ± 24% vs. 313% ± 103%). The study noted that to apply magnetic targeting in humans, due to anatomical differences, a superconducting magnetic coil or more strongly magnetic nanoparticles might be necessary [73]. Similar findings were observed with BCNU [74]. Both MRgFUS alone and magnetic targeting alone doubled magnetic nanoparticle concentration in the treated brain regions compared to untreated regions, while the combination of magnetic nanoparticles, MRgFUS, and magnetic targeting increased magnetic nanoparticles accumulation by nearly 10-fold compared to untreated regions and 26-fold compared to magnetic nanoparticles without additional treatment. A week after treatment, the medium-dose magnetic nanoparticles, MRgFUS, and magnetic targeting group showed significantly delayed tumor progression compared to no treatment (-0.79 ± 0.35 cm ³ vs. 2.98 ± 2.61 cm ³ ) or magnetic nnaoparticles alone (-0.79 ± 0.35 cm ³ vs . 2.96 ± 3.00 cm ³ ). Furthermore, magnetic nanoparticles alone were more effective than unbound BCNU (1.15 ± 1.58 cm ³ vs. 2.48 ± 3.09 cm ³ ), with the effect being dose-dependent [74].

To simplify protocols involving magnetic nanoparticles, researchers have developed microbubbles conjugated with superparamagnetic iron oxide nanoparticles (SPIONs), a type of magnetic nanoparticles, loaded with DOX [75]. This single formulation combines the functions of microbubbles, gadolinium contrast, and chemotherapy-conjugated magnetic nanoparticles, enhanced by MRgFUS and magnetic targeting. In a rat glioma model, SPION deposition increased fourfold with MRgFUS and magnetic targeting compared to the non-targeted hemisphere, while MRgFUS alone and magnetic targeting alone resulted in 2.7-fold and 2.3-fold increases, respectively. Additionally, DOX deposition in the treated hemisphere increased twofold with MRgFUS and magnetic targeting compared to the control. A follow-up study using an improved formulation that enhanced DOX-carrying capacity and R2 relaxivity demonstrated that SPION-DOX complexes accumulated 2.8 times more in MRgFUS-targeted brain tissue with MT compared to without magnetic targeting. DOX deposition was also enhanced by more than 2.1 times with magnetic targeting. The MR R2 value was highly correlated with SPION concentration (R ² = 0.83) and DOX accumulation (R ² = 0.79), suggesting that SPION-DOX microbubbles may be advantageous for dosing monitoring through imaging [75].

Advancements in neuro-oncology have been impeded by the challenge of delivering therapeutics across the BBB [76]. FUS has emerged as a key technique in this field, enhancing the delivery of both small and large molecular therapeutics to the CNS (Table 1). For instance, MRgFUS has facilitated the transport of molecules such as BCNU [77], cisplatin [78], and doxorubicin [79,80], as well as larger molecules like trastuzumab [81] and adeno-associated viruses [82]. In a pivotal first-in-human study, MRgFUS was demonstrated to be safe for opening the BBB in patients with high-grade gliomas, setting the stage for further clinical trials [1,83].

Subsequent studies, including a multi-center trial, have explored the safety and efficacy of MRgFUS in conjunction with maintenance temozolomide therapy for newly diagnosed glioblastoma multiforme, with results pending publication (NCT03616860, accessed 1 March 2024) [1]. Moreover, the technique's application in enhancing the delivery of radio-labeled therapeutics, such as Indium-111-radiolabeled trastuzumab in patients with Her2-positive breast cancer metastases, has shown promise in reducing tumor volume [84]). Ongoing research, including trials examining the enhanced delivery of pembrolizumab for brain metastases from non-small cell lung cancer, continues to refine the understanding of MRgFUS in clinical practice (NCT05317858). These studies underscore the critical role of pharmacokinetic data and imaging biomarkers in optimizing the therapeutic use of MRgFUS, despite existing gaps in our knowledge [1,83-88] (Table 1).

MRgFUS interventions for Alzheimer’s disease

Alzheimer's disease (AD) is a progressive neurodegenerative disorder and the leading cause of dementia. Its pathology involves the buildup of toxic amyloid-beta (Aβ) plaques outside neurons, tangles of tau protein inside neurons, neuronal loss, particularly in the circuit of Papez, and dysfunction in the default mode network (DMN) [89-91]. The formation of Aβ plaques is accelerated throughout the brains of AD patients, particularly in areas vulnerable along the DMN, correlating with cognitive decline [92,93]. There is significant interest in using targeted therapies like monoclonal antibodies, including aducanumab and lecanemab, to reduce Aβ plaques in the central nervous system [94]. The FDA approved aducanumab in 2021 based solely on its ability to clear Aβ [95], and lecanemab received approval in 2023 following evidence from a randomized controlled trial showing both Aβ reduction and a modest slowing of cognitive decline [96]. Despite these developments, only a tiny fraction (0.01%) of Aβ antibodies cross the BBB, due to their large molecular size, which the BBB typically filters out, highlighting the need for alternative delivery methods like ultrasound (Table 2) [97,98].

In 2018, the first clinical application of MRgFUS to safely open the BBB in AD patients was reported [59]. This initial study aimed primarily to establish safety, involved opening the BBB in a small region (approximately 9 × 9 × 9 mm) twice, a month apart [59]. T1 gadolinium-enhanced MRI confirmed that the BBB opening was temporary, closing within 24 hours [59]. Subsequent studies expanded on this, demonstrating safe BBB opening over larger volumes and targeting different anatomical sites [99,100]. More recently, MRgFUS was used to open the BBB at major DMN nodes, including the bilateral hippocampi and precuneus, to facilitate drug delivery to these areas [60].

Preclinical studies indicated that FUS could increase aducanumab's penetration into the CNS by 5-8 times [101]. Building on these findings, a recent first-in-human trial utilized MRgFUS to enhance the delivery of aducanumab in three AD patients [102]. The treatment involved six monthly MRgFUS-mediated BBB openings paired with increasing doses of intravenous aducanumab. The amount of brain tissue targeted increased across patients, from 10 mL in the non-dominant frontal lobe of the first patient to 40 mL in the dominant frontal, temporal, and hippocampal regions of the third patient. Aβ reduction was observed in the treated areas using fluorine-18 florbetaben positron emission tomography (PET) scans, with untreated homologous brain regions serving as controls [102]. However, the third patient experienced cognitive decline by the 30-day follow-up, raising challenges in distinguishing between natural disease progression and potential adverse effects of the treatment. Additionally, there remains uncertainty about whether reducing Aβ levels directly correlates with cognitive improvement [103] (Table 2).

Therapeutic innovations for Parkinson’s disease using MRgFUS

Parkinson's disease (PD) is the most prevalent neurodegenerative movement disorder, affecting 1% of individuals over the age of 60 [106]. The hallmark motor symptoms of PD are due to the loss of dopaminergic neurons in the substantia nigra pars compacta, accompanied by Lewy body pathology, resulting in reduced dopamine levels in the striatum [107]. The exact causes of PD are still being explored, but they likely involve a combination of lifestyle factors, environmental exposures, and genetic predispositions [108]. While aging is the primary risk factor, other contributors include exposure to toxins like rotenone and paraquat [109] and genetic mutations such as those in the GBA1 [110,111] and LRRK2 genes [112]. As in Alzheimer's disease and neuro-oncology, BBB has been a significant challenge for PD therapies, sparking interest in using FUS to enhance the delivery of treatments targeting alpha-synuclein or neurotrophic factors to the basal ganglia [113].

In preclinical models, studies have shown that the BBB can be safely opened in the putamen [114]. When combined with alpha-synuclein silencing viral vectors or neurotrophic factors, this approach can reduce nigrostriatal neuron loss in MPTP or transgenic mouse models [115,116]. A recent study highlighted the use of MRgFUS to open the BBB, combined with systemic administration of an AAV vector, inducing novel protein expression in the putamen and substantia nigra, which may offer a new strategy for protein modulation in PD patients [61]. However, challenges include the high cost of systemic AAV delivery at effective doses for human transduction and potential risks from systemic exposure, even with improved BBB permeability. Advances in neuron-selective AAV serotypes with minimal systemic uptake may provide a solution for integrating gene therapy with MRgFUS [117].

In clinical settings, MRgFUS-mediated BBB opening in the putamen has shown good tolerance in PD patients, including in bilateral and repeated treatments [91]. The putamen is sensitive to physical damage [118], yet MRgFUS treatments have not shown exacerbation of dopaminergic denervation in imaging studies, although data remain limited and further research is needed to define safety thresholds [119]. The first clinical use of MRgFUS-mediated BBB opening paired with drug delivery was recently reported in PD patients with GBA1 mutations, which are linked to Gaucher's disease [120]. GBA1 encodes the enzyme glucocerebrosidase (GCase), and preclinical studies suggest that a deficiency or reduced activity of GCase leads to alpha-synuclein accumulation, which may contribute to dopaminergic neuron loss [121,122]. Intravenous recombinant GCase is a promising therapy for PD in the context of GBA1 mutations, though its penetration of the BBB is limited [123].

A study was conducted to evaluate the safety of combining MRgFUS-mediated BBB opening with intravenous GCase administration in patients with GBA1-related PD [124]. Four patients underwent a total of three MRgFUS procedures, each followed by an increasing dose of GCase, spaced two weeks apart. Gadolinium-enhanced MRI confirmed successful unilateral BBB opening in the putamen without severe adverse effects [120]. Two patients reported transient increases in dyskinesia, possibly due to increased levodopa exposure from BBB permeabilization [120]. Positron emission tomography (PET) scans showed reduced metabolism in the putamen one-month post-treatment, correlating with treatment efficacy seen in other studies [125]. This phase 1 trial lays the groundwork for targeted delivery of GCase and other therapeutic molecules, such as neurotrophins or monoclonal antibodies, for movement disorders. Future research will focus on understanding the safety of MRgFUS in PD patients and the pharmacodynamics of GCase when delivered via ultrasound.

Advanced Neurological Therapies with MRgFUS

MRgFUS-enhanced immunotherapy

Research on the use of immunotherapy combined with MRgFUS for treating brain tumors has primarily focused on delivering HER2-targeting antibodies for breast cancer metastases. In a mouse model, FUS successfully delivered trastuzumab, although the dosage was limited due to red blood cell extravasation [81]. Without MRgFUS, trastuzumab levels in brain tissue were undetectable except in one case. In a rat model of HER2-positive human breast cancer brain metastases, Park, et al. demonstrated that MRgFUS-enhanced delivery of trastuzumab extended the median survival time compared to no treatment (83 days vs . 63 days) and trastuzumab alone (83 days vs. 71 days, not statistically significant), with a significant reduction in tumor volume at week 7 [126]. Notably, this effect was observed in a subset of the experimental group, as 6 of the 10 treated mice did not respond. Rats treated with MRgFUS and HER2-specific NK-92 cells administered peripherally showed increased mean survival times [127]. Similar to Park, et al.'s findings, approximately half of the treated animals were non-responders, exhibiting survival curves similar to untreated animals. The study did not observe histological evidence of red blood cell extravasation, possibly because there was a delay of a week or more between the last MRgFUS treatment and euthanasia.

In a rat glioma model, FUS increased the concentration of intraperitoneal IL-12 in the brain nearly 2.9-fold compared to no FUS [128]. The combination of IL-12 and FUS significantly increased the presence of T-lymphocytes within the tumors compared to sham treatment (CD3+ CD4+: more than 4-fold increase; CD3+ CD8+: 5-fold increase; CD4+ CD25+: 2-fold increase; cytotoxic-to-regulatory T-cell ratio increase: 2.5-fold). These effects were not observed in healthy rats or systemically in tumor-bearing rats. The median survival time was also longer compared to no treatment (30 days vs . 21 days) and IL-12 alone (30 days vs . 26 days). The promising results in animal models suggest that immunotherapies could potentially be integrated into MRgFUS treatment protocols, pending further research to optimize targeting strategies [128].

Gene therapy augmentation with MRgFUS

The use FUS to deliver DNA-loaded microbubbles represents a novel approach in the field. In a rat glioma model, folate-conjugated cationic microbubbles (cMBs) containing DNA were used in conjunction with ultrasound targeting, resulting in successful transfection of tumor cells. This treatment increased reporter gene expression in tumor tissue by 4.7 times compared to direct injection and by 1.5 times compared to cMBs without folate conjugation, with expression localized exclusively within the tumor [129]. Another study in a rat glioma model used DNA-loaded cMBs conjugated with VEGFR2-targeted monoclonal antibodies. This approach resulted in reporter gene expression that was 3.7 times higher with targeted cMBs and 2.3 times higher with non-targeted cMBs compared to direct DNA injection [130]. Additionally, the study explored the use of cMBs with the suicide gene pHSV-TK, which converts ganciclovir into a compound that halts DNA replication, leading to tumor cell death [131]. The combination of cMBs and FUS significantly reduced tumor volume at day 25 compared to direct injection (9.7 ± 5.2 mm ³ vs . 40.1 ± 4.3 mm ³ ) and non-targeted cMBs (9.7 ± 5.2 mm ³ vs . 21.8 ± 4.7 mm ³ ). Both targeted and non-targeted cMBs led to a significant reduction in tumor volume and an increase in median survival time compared to untreated controls [131].

Gene therapy, among brain tumor treatments using FUS, remains the least developed. As this technology progresses towards clinical trials, incorporating MR guidance will be crucial for imaging through the human skull.

Challenges and Considerations of MRgFUS

While preclinical studies in animal models have shown promise, comprehensive testing in humans is necessary. A Phase 1 clinical trial is essential to assess toxicity and establish safe parameters for the strength of MRgFUS, as well as to optimize the selection, dosage, and timing of systemic chemotherapeutic agents. The human skull is thicker than that of rodents, and the greater distances within the human cranium can reduce ultrasound intensity, necessitating higher power levels. The depth of tumors may also restrict the types of lesions that can be effectively treated [30]. Moreover, severe or symptomatic peritumoral edema might worsen with further disruption of the BTB. Using stereotactic frames poses challenges for patients needing multiple FUS treatments, but frameless systems are in development [31]. The risk of hemorrhage has been observed in animal models, but it remains unclear how this will manifest in human patients [67,81].

Conclusion

MRgFUS offers a significant advancement in non-invasive, targeted drug delivery for brain tumors. By combining MRgFUS with other technologies such as nanoparticles, the precision and effectiveness of treatments can be significantly enhanced. Preclinical studies have shown that MRgFUS improves the delivery and efficacy of chemotherapy, immunotherapy, and gene therapy, leading to better treatment outcomes, reduced tumor growth, and extended survival times.

Looking ahead, there is growing interest in applying MRgFUS-mediated BBB opening to a broader range of neurological conditions, including Alzheimer's disease, Parkinson's disease, and neuro-oncological disorders. As clinical trials expand and involve larger cohorts, a more comprehensive understanding of the biological effects of drugs delivered via MRgFUS will emerge, allowing for more precise dosing and optimized treatment schedules. Randomized controlled trials will likely provide more detailed data on the efficacy of this method across different indications.

Technological advancements in MRgFUS devices, such as the development of frameless systems and minimally invasive procedures, are anticipated to further enhance the method's accessibility and patient comfort. These innovations will be essential for integrating MRgFUS into standard therapeutic protocols, potentially transforming the treatment landscape for patients with complex neurological conditions.

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