Systems and methods for modulation of deep brain circuits

Abstract

A system and method for applying therapeutic ultrasound to the brain while using ultrasound to compensate for ultrasound attenuation and phase dispersion by an individual's head. The compensation delivers a deterministic ultrasound intensity to the target. The compensation is based on relative ultrasound transmission measurements, which are made using a set of ultrasound emitters on one side of the head and a set of receivers on the other side. Measurements are made without and with the head. Based on the difference between these measurements, the set of ultrasound waves is adjusted to compensate for the attenuation and phase dispersion caused by ultrasound passing through the skull and scalp into the head. The adjusted set of ultrasound waves delivers a deterministic target ultrasound intensity to the target location. This deterministic delivery enables safe and effective ultrasound neuromodulation, safe and effective localized drug release from nanoparticle carriers, and safe and effective blood-brain barrier disruption using microbubbles to deliver drugs, genes and stem cells across the blood-brain barrier.

Description

[Technical field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application is a nonprovisional application claiming priority to U.S. Provisional Patent Application No. 63 / 296,252, filed January 4, 2022, the contents of which are incorporated herein by reference.

[0002]

[0002] The present disclosure relates to systems and methods for applying and modulating ultrasound to specific targets within the human brain. [Background technology]

[0003] Approximately one-third of patients with psychiatric and neurological disorders are treatment-resistant. Neuromodulation has the potential to target and reset dysfunctional circuits. However, current neuromodulation approaches are limited by two main barriers. First, the exact neural circuits involved in psychiatric and neurological disorders are poorly understood and appear to vary among individuals. This uncertainty limits the reliable application of deep brain stimulation to such patients. Second, common disorders such as depression, anxiety, and pain-related disorders involve neural networks located deep in the brain, including limbic, basal ganglia, and brainstem networks. These deep brain sources complicate treatment using current noninvasive neuromodulation approaches. For example, electroconvulsive therapy modulates deep brain structures using large currents that induce brain-wide seizures. Such widespread activation often leads to cognitive side effects such as memory loss. Transcranial magnetic stimulation may be able to modulate deep brain circuits through connections with stimulated cortical areas, but such indirect effects contribute to response variability.

[0004]

[0004] Ultrasound-based neuromodulation has the potential to selectively modulate targets deep in the brain with high spatiotemporal resolution.

[0005] Compared to studies in small animal models, neuromodulation using ultrasound has shown limited efficacy when applied to the human head. It has been found that the intensity of neuromodulation ultrasound is attenuated by a factor of 4.5 to 64 in the human skull alone, depending on the skull segment and individual differences. This very large variability in attenuation coefficients makes it impossible to make a reliable estimate of the intensity delivered. As a result, human studies are forced to take a "worst case scenario" approach, assuming as little attenuation as possible, to reduce the risk of harming the brain. This conservative approach, necessary for safety, has limited efficacy.

[0005]

[0006] Furthermore, effective and safe treatment of brain disorders requires selective delivery of ultrasound to localized brain regions or individual nuclei. Current single-element transducers used for neuromodulation typically only form cigar-shaped beams spanning a few centimeters, limiting spatial specificity. In addition, these solutions lack the precision and flexibility of phased arrays, which complicates selective delivery, fine-tuning of targeting, and systematic application to multiple targets. The lack of precision and flexibility limits the use of existing solutions for patients with deep brain circuit disorders.

[0006]

[0007] Therefore, a device that compensates for the human skull and delivers ultrasound to specific deep brain targets with high spatiotemporal resolution is desired. Summary of the Invention [Means for solving the problem]

[0007]

[0008] Transcranial focused ultrasound offers a noninvasive and reversible approach to precisely and individually manipulate brain circuits, potentially transforming our understanding of brain function and the treatment of brain dysfunction. Ultrasound can be focused through the intact skull and scalp to specific deep brain regions that can span several millimeters in diameter.

[0008]

[0009] Because ultrasound reaches its target in microseconds, ultrasound arrays have the ability to stimulate multiple sites simultaneously or in precise time sequences. Precise focusing on command opens up unique new possibilities for systematically modulating dysfunctional circuits in each individual. Furthermore, the transducer array can be programmatically focused on specific brain targets without moving the device or the subject.

[0009]

[0010] However, as mentioned above, the efficacy and safety of these approaches have been limited by the human head, which strongly and unpredictably attenuates and distorts ultrasound waves. This barrier is due to the strong and unpredictable attenuation of ultrasound waves by the head and the limited targeting accuracy of existing devices. To address these issues, ultrasound phased array devices have been developed that compensate for the distortion of ultrasound waves by the head and deliver ultrasound waves to specific targets with high spatiotemporal resolution. The devices can be used to confirm the action on the target inside the MRI and can be used repeatedly outside the MRI.

[0010]

[0011] To address the issue of ultrasound attenuation, a "Relative Through-Transmit" (RTT) approach was developed that directly measures and compensates for attenuation and distortions of a given skull and scalp. RTT was implemented in hardware and demonstrated to accurately restore the operator's target intensity in an ex vivo human skull. Furthermore, this functionality enabled effective, intensity-dependent transcranial modulation of nerves and the effective delivery of a prescribed dose of propofol into the skull. Thus, the present disclosure provides a tool to non-invasively and effectively modulate specific neural circuits deep in the human brain, providing a treatment option for millions of people who are resistant to current therapies.

[0011]

[0012] The devices disclosed herein provide diagnostic information to guide deep brain stimulation implants, improve understanding of human brain function, and offer new means to induce durable circuit resetting in treatment-resistant patients.

[0012]

[0013] In some embodiments, the systems and methods described herein provide precision ultrasound therapy that can be used, for example, in mental health or neurological clinics. In some embodiments, the system provides a mechanism for controlling the amount of ultrasound delivered to the head to produce predictable effects in target areas. In some embodiments, the system allows for multi-focal manipulation to predictably modulate specific brain regions based on the specific needs of the patient.

[0013]

[0014] In some embodiments, the system comprises a head-mounted transducer array device with 64-1,024 transducer elements inserted into a 3D printed frame with a geometry optimized for a particular multifocal operation in a particular individual. Ultrasound delivery to deep brain targets is possible because of minimal attenuation of ultrasound by brain tissue. However, the head and skull in particular dephase and attenuate ultrasound. Thus, in some embodiments, the systems and methods described herein use the ultrasound itself to correct (also called "compensate") for aberrations of ultrasound caused by the head. In this way, the aberrations of the head are measured directly and accurately, without the need for further head scans such as CT or MRI. In particular, the method performs ultrasound RTT measurements of each corresponding segment of a given head. This results in phase and amplitude values ​​that are used to correct the aberrations of each segment of a particular head. The amplitude of each ultrasound transducer is scaled and phase shifted, for example, to deliver undistorted, deterministic intensity to the treatment target. In some embodiments, certain characteristics of the transmission waveform may be optimized to maximize the accuracy of detection of, and therefore correction of, the ultrasonic energy delivered through the head.

[0014]

[0015] In some embodiments, the systems and methods described herein are configured and adapted for the treatment of anxiety disorders and depression-related disorders, including post-traumatic stress disorder. These disorders involve abnormal connections between two deep brain regions, the subgenual cingulate cortex and the amygdala, and adjacent circuits. Low-intensity ultrasound for tens of seconds targeting the cingulate cortex and the amygdala can induce lasting changes in the associated circuits. In some embodiments, the systems and methods described herein are configured to target these regions (e.g., using human cadavers as models).

[0015]

[0016] In some embodiments, the systems and methods described herein are configured and adapted for the treatment of thalamic nuclei that are involved in pain. These disorders involve abnormal connections of the insular cortex, cingulate cortex, nucleus accumbens, and ventral tegmental area of ​​the thalamic nuclei. Tens of seconds of ultrasound targeting these circuits modulates pain thresholds. In some embodiments, the systems and methods described herein are configured to target these regions (e.g., using human cadavers as models).

[0016]

[0017] In some embodiments, targeting of these regions can be confirmed using fMRI BOLD (blood oxygen level dependent imaging), MRI thermometry, or MRI acoustic radiation force imaging. These imaging sequences visualize the areas affected by ultrasound, thus increasing the reproducibility of ultrasound therapy and minimizing potential off-target effects.

[0017]

[0018] In some embodiments, MRI can be used to establish subject-specific anatomy of the head and brain.

[0019] In one embodiment, the present disclosure provides a method for applying transcranial ultrasound to a target brain location. At least one transmitting ultrasound transducer is driven to generate ultrasound waves and achieve a target ultrasound energy at a target location within a free-field volume corresponding to the target brain location. The ultrasound waves exiting the free-field volume are measured by at least one receiving ultrasound transducer located at a fixed distance and orientation relative to the at least one transmitting ultrasound transducer on the opposite side of the free-field volume. The head is then placed between the transmitting and receiving ultrasound transducers, the transmitting ultrasound transducer is again driven to generate the same ultrasound waves into the head, and the receiving ultrasound transducer measures the ultrasound waves exiting the head, which have been altered at least in part due to the presence of the head in the ultrasound path between the transmitting and receiving ultrasound transducers. One or more adjusted ultrasound waveforms are then determined based on the difference between the measured ultrasound passing through the free-field volume and the measured altered ultrasound passing through the head, where the adjusted ultrasound compensates for attenuation and phase shift due to obstacles in the ultrasound path to deliver actual ultrasound stimulation energy to the target brain location so as to approximate the target ultrasound stimulation energy at the target brain location. A transmitting ultrasound transducer is then driven to generate the adjusted ultrasound into the head.

[0018]

[0020] In another embodiment, the present disclosure provides an ultrasound-based neurostimulation system comprising a head-mounted device and a controller. The head-mounted device comprises at least one array of ultrasound transducers, including ultrasound transducers disposed on either side of the head. The controller is configured to determine a set of ultrasound waves to be transmitted by the ultrasound transducers to achieve a target ultrasound stimulation energy at a target location within the volume. The controller drives a first set of ultrasound transducers and captures the propagating ultrasound waves using a second set of ultrasound transducers on the opposite side of the volume. This is done twice: once in a free field while the head-mounted device is not applied to the head, and once while the head-mounted device is applied to the head. The controller compares the free field measurements with the transmission measurements to determine attenuation and phase shift of the ultrasound waves at least partially due to the presence of the skull in the ultrasound path. The controller determines an adjusted set of ultrasound waves that compensates for the determined attenuation and determined phase shift to achieve a target ultrasound stimulation energy at a target location within the volume while the head-mounted device is applied to the head. The controller drives both sets of transducers according to the determined coordinated set of ultrasound waves while the head-mounted device is applied to the head.

[0019]

[0021] In yet another embodiment, the present disclosure provides a method for applying a deterministic ultrasound dose to a target brain location by driving an ultrasound transducer array to perform through-the-head ultrasound measurements and without the subject's head. Based on the through-the-head ultrasound measurements, ultrasound attenuation and phase shift due to the head compared to without the head are determined. Corrected amplitude and phase values ​​are determined for each ultrasound transducer, and the ultrasound transducer array is driven based on the adjusted amplitude and phase values ​​to achieve a target ultrasound stimulation energy at the intracranial target brain location.

[0020]

[0022] Other aspects of the invention will become apparent by consideration of the detailed description and accompanying drawings.

[0023] The patent or application file contains at least one color drawing. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee. [Brief description of the drawings]

[0021] [Figure 1]

[0024] FIG. 1 is a block diagram of a system for applying ultrasound stimulation according to one embodiment. [Diagram 2]

[0025] FIG. 2 is a perspective view of an example of a head-mounted transducer array device of the system of FIG. 1. [Diagram 3]

[0026] FIG. 2 is a perspective view of another example of a head-mounted transducer array device of the system of FIG. 1. [Figure 4]

[0027] 4 is a series of MRI images taken of a subject wearing the head-mounted transducer array device of FIG. 3 and a subject not wearing the head-mounted transducer array device. [Figure 5A]

[0028] FIG. 4 is a perspective view of the transducer array of the head-mounted transducer array device of FIG. 3 transmitting ultrasound waves through a free-field volume. [Figure 5B]

[0029] FIG. 4 is a perspective view of the transducer array of the head-mounted transducer array device of FIG. 3 transmitting ultrasound waves through the head of a subject. [Figure 5C]

[0030] FIG. 1 is a diagram of the superposition of ultrasound beams from a transducer array onto an MRI image. [Figure 6]

[0031] 1 is a graph of intensity attenuation through individual segments of the skull. [Figure 7]

[0032] 5D is a graph of the target intensity and actual intensity of transmitted ultrasound at the target location due to attenuation in the ultrasound target shown in FIG. 5C. [Figure 8]

[0033] 1A-1C are schematic diagrams illustrating a method for measuring ultrasound aberration due to the skull, according to one embodiment of the present disclosure. [Figure 9]

[0034] 1 is a flow chart of a method of applying ultrasound stimulation to a target location within a subject's brain after adjusting the stimulation wave to compensate for attenuation and phase shifts caused by the subject's skull. [Figure 10]

[0035] 10 is a series of graphs showing the measured intensity of a transmitted ultrasound beam in the absence of a skull, the measured intensity of the same transmitted ultrasound beam attenuated by a human skull (n=8), and the measured intensity of the ultrasound beam passing through the skull after correction by the method of Fig. 9. The bars show how the method of Fig. 9 recovers the target intensity for deep brain targets. [Figure 11]

[0036] 10 is a graph showing the correction accuracy of the method of FIG. 9 for stimulation of the nerve of a subject's thumb placed inside the skull. [Figure 12]

[0037] 12 is a graph showing subject response data to the highest ultrasound pressure shown in FIG. 11 separately for when the skull is not present, when the skull is present, when the skull is present and the ultrasound is compensated according to the method of FIG. 9, and for a stimulus offset 10 mm from the target. [Figure 13]

[0038] 10 is a graph showing measured peak pressures for a target stimulus, an actual stimulus (after cranial attenuation), and a stimulus compensated according to the method of FIG. 9. [Figure 14]

[0039] 10 is a graph showing the variation in focal volume of the area stimulated by applied ultrasound due to distortion by the skull and compensation of the ultrasound by the method of FIG. 9. [Figure 15]

[0040] FIG. 2 is a perspective view of an example of a head-mounted transducer array device of the system of FIG. 1 applied to a patient. [Figure 16]

[0041] 1 shows that a head-mounted transducer array device produces a spatially focused intensity field through the skull.10 Intensity field of the device's focal point measured through a human skull ex vivo. [Figure 17A]

[0042] FIG. 1 shows the intensity field generated by the transducer array superimposed on the subject's brain anatomy for scale. [Figure 17B] FIG. 1 shows the intensity field generated by the transducer array superimposed on the subject's brain anatomy for scale. [Figure 17C]

[0043] FIG. 1 shows fMRI BOLD responses produced by stimulation in target regions. [Figure 17D] FIG. 1 shows fMRI BOLD responses produced by stimulation in target regions. [Figure 17E]

[0044] FIG. 1 shows that fMRI BOLD responses at targets are time-locked to stimulus onset. [Figure 17F]

[0045] FIG. 13 shows active sham stimulation with the same stimulation parameters, but with an array (plane wave) focused outside the subject's head, did not evoke a BOLD response at the target. [Figure 17G] FIG. 13 shows active sham stimulation with the same stimulation parameters, but with an array (plane wave) focused outside the subject's head, did not evoke a BOLD response at the target. [Figure 18]

[0046] FIG. 1 shows in chart form that stimulation improves mood state in depressed patients. Self-reported mood scores after each stimulation indicate a reduction in anxiety and depression and an increase in well-being. Sham stimulation did not induce any changes in the subjects' mood ratings. [Figure 19]

[0047] Figure 1 shows an exemplary correction method that uses transmission measurements from every element pair to resolve the attenuation and speed-up that occurs in front of each element, and then adjusts the stimulation parameters to correct for these distortions. [Figure 20]

[0048] Figure 1 shows the average attenuation correction across sessions. Variability was observed across sessions due to the quality of acoustic coupling between the transducer and the subject's head. Elements that were not properly coupled (mean pressure transmissibility below 12%) were switched off. [Figure 21]

[0049] 13 is a chart showing that mechanical resist provides reproducible targeting across sessions. a. Deviation between the location of the ultrasound focus and the brain target for five human subjects. b. Targeting error represented by spatial dimensions. [Figure 22]

[0050] 13 is a chart showing that mechanical resist provides reproducible targeting across sessions. a. Deviation between the location of the ultrasound focus and the brain target for five human subjects. b. Targeting error represented by spatial dimensions. [Diagram 23]

[0051] The blue region is an image showing the electronic steering range available to a head-mounted transducer array device. The focal spot of the device through an ex-vivo human skull is superimposed on the MRI image for scale. The focal spot can be beamformed to any position within this region in microseconds without physical movement of the device. [Figure 24]

[0052] FIG. 1 shows ultrasound attenuation in the human skull. [Diagram 25]

[0053] Figure 1 shows that RTT accurately compensates for each skull and recovers the target intensity at the target. Ultrasound fields obtained in ex vivo human skulls (n=8) with hypothetical ideal compensation (grey), no compensation (red), and RTT (green). The top bar indicates the spatial peak intensity of the ultrasound field in each case. The bottom plot shows the corresponding spatial distribution of the ultrasound field relative to the target. [Figure 26]

[0054] Figure 14. Effect of RTT applied to human head. Average transmission attenuation values ​​across all elements in five human subjects and eight ex-vivo human skulls. No shaving was required to obtain a robust transmission signal. [Figure 27]

[0055] Figure 1 shows the RTT performance as a function of working range, including targets that make up the full working range of the array: 10mm axial, 20mm axial, 10mm lateral, 20mm lateral, and 15mm elevation relative to the central target. Axial refers to the line joining the centers of the two transducers. [Figure 28]

[0056] Phase correction is insufficient to account for cranial attenuation. Spatial peak intensity at the central target using the phase-corrected (a), amplitude (b) and both (c) components of the RTT. [Figure 29]

[0057] Figure 2: RTT is robust regardless of hardware. (a) Array geometry of a head-mounted transducer array device. Associated correction values ​​are the same data as in Fig. 25. Subject 8 (purple marker) with the thickest skull due to bony prominences and possibly hyperostosis was relatively poorly corrected. (b) Array geometry of a head-mounted transducer array device with a larger aperture and associated correction data. [Diagram 30]

[0058] Figure 1. RTT enables effective ultrasound stimulation through the skull. (a) RTT enables effective modulation of peripheral nerves through the skull. Arrays were targeted to the nerves of the thumbs of 11 participants. The thumbs were fixed to a central target in an ex vivo skull. The arrays delivered 300 ms stimuli to the targets with a frequency of 650 kHz and a pressure amplitude of 1.8 MPa. Data were collected with ideal compensation (black), without compensation (red), and after RTT application (green). The sham condition delivered stimuli 10 mm below the finger (yellow). Individual conditions were presented randomly every 8–12 s and repeated a total of 10 times. Subjects reported nociceptive responses indicative of stimulation to the nerve and nerve endings. Response frequency indicates the percentage of trials in which subjects reported a nociceptive response. (b) Relationship between stimulation dose and response. Responses were significantly modulated by ultrasound pressure, but did not differ significantly between ideal (black) and RTT (green) corrected responses (see text for details). Low, medium, and high labels correspond to target peak pressures of 1.3 MPa, 1.55 MPa, and 1.8 MPa measured in the free field. Error bars represent s.e.m. [Diagram 31]

[0059] Figure 1. Ideal-corrected (black) and RTT-corrected (green) mean response rates for each subject. When RTT was not applied, there was no significant ultrasound nerve stimulation (red; t11=1.00, p=0.34, one-sample t-test). [Diagram 32]

[0060] Figure 3 shows that RTT enables effective and dose-dependent local drug release. (a) RTT enables effective drug release from nanoparticle carriers. The safe and biocompatible nanoparticle carrier was loaded with propofol, a neuromodulation drug. The nanoparticles release their drug load when bombarded with low-to-moderate intensity ultrasound. The vial containing the nanoparticles was placed in a central position on the ex vivo skull, similar to Figure 30. Ultrasound was delivered through the skull with a frequency of 650 kHz and a pressure amplitude of 1.8 MPa, with 100 ms pulses delivered every 1 s for 60 seconds to bombard the nanoparticles. Data were collected under ideal compensation of the skull assumption (black), without compensation (red), and after application of RTT (green). The sham condition delivered the stimulus 10 mm below the vial (yellow). In a second sham condition, the vial was placed on the target, but no ultrasound was delivered (purple). The dotted line represents the baseline when no ultrasound was applied. The baseline may not be zero due to partial leakage of free (unencapsulated) drug or nanoparticles. As in Figure 30, individual conditions were randomly interleaved. Bars consist of n=10 separate samples, except for no ultrasound, where n=6 was used. Error bars represent s.e.m. (b) Dose-response relationships. Low, medium, and high labels correspond to target peak pressures of 1.2 MPa, 1.5 MPa, and 1.8 MPa measured in the free field. All data points consist of n=10 separate samples. Error bars represent s.e.m. [Diagram 33]

[0061] Compensation for ultrasound attenuation by the head is important for effective neuromodulation: the figure shows the statistical significance of fMRI BOLD signal modulation by ultrasound in deep brain targets when compensation is applied (left) and not (right). DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0022]

[0062] Before describing embodiments of the invention in detail, it is to be understood that the invention is not limited in its application to the details of construction and the arrangement of components set forth in the following description or illustrated in the following drawings. The invention is capable of other embodiments and of being practiced or carried out in various ways.

[0023]

[0063] The systems and methods illustrated in the following examples provide a non-invasive ultrasound-based neuromodulation device that may be used, for example, to treat and / or diagnose neurological and psychiatric disorders. The device is worn on the head and coupled at several locations around the skull. This arrangement allows the device to deliver ultrasound from the transducers to designated deep brain targets. The device is controlled by an operator who can select specific brain regions to target with ultrasound. A patient treatment plan may be created to automatically control which transducers deliver ultrasound to the deep brain targets. The device typically performs compensation procedures to compensate for ultrasound obstructions, including the skull, hair, and ultrasound coupling, and emits ultrasound from the individual transducer arrays at amplitudes and timings such that the ultrasound reaches the predefined targets at the target intensity.

[0024]

[0064] In some embodiments, the system is configured to correct ultrasonic aberrations in the skull by using actual ultrasound waves of the same frequency as those used for treatment. The device directly measures the skull as opposed to estimates, allowing for more accurate, safe and effective treatment. The device obtains this measurement by measuring the actual attenuation and phase dispersion of the ultrasound waves via relative transmission measurements through the skull. The system then takes these measured attenuation and phase values ​​and compensates for them by adjusting the amplitude and phase of each element. Measuring ultrasonic cranial aberrations using actual ultrasound allows for precise compensation of the ultrasound intensity delivered to specific brain targets, resulting in sharper ultrasound focus, resulting in more accurate, safe and effective treatment.

[0025]

[0065] In some implementations, the head-mounted device includes 256 transducer elements mounted on a patient-specific, custom 3D printed frame. This modular design allows for targeting of specific brain regions for each patient.

[0026]

[0066] FIG. 1 illustrates an example of an ultrasound-based neurostimulation system 100 for deep brain therapy. The system 100 includes a controller 101 with an electronic processor 103 and a non-transitory computer-readable memory 105. The electronic processor 103 is communicatively connected to the memory 105 and configured to store data in and access stored data from the memory 105. The memory 105 also stores computer-executable instructions that, when executed by the electronic processor 103, provide the functionality of the controller 101, including, for example, functionality described herein. While the example of FIG. 1 illustrates only one memory 105 in other embodiments, the system may utilize multiple different memory modules, including, for example, local memory, external storage devices, and / or remote or cloud-based memory systems. Similarly, in different embodiments, the system 100 may utilize one or more electronic processors implemented in one or more different computing devices. In some embodiments, the controller 101 may be implemented as an application-specific controller device, and in some embodiments, the controller 101 may be provided as a desktop, laptop, or tablet computer. In yet other implementations, the controller 101 may comprise a number of different control devices, including, for example, an electronic controller integrated into or directly connected to a head-mounted device, as described below, and a computer communicatively connected to the electronic controller. Thus, unless otherwise noted, the controller 101 may comprise one or more computing devices and / or control circuits, one or more electronic processors, and one or more memories.

[0027]

[0067] 1, the controller 101 is communicatively connected to a plurality of ultrasonic transducers 107, including ultrasonic transducers 107.1, 107.2, and 107.n. In the illustrative embodiment described herein, the plurality of ultrasonic transducers 107 includes 256 ultrasonic transducers arranged in one or more arrays, which are incorporated into a head-mounted device.

[0028]

[0068] As described in further detail below, the controller 101 is configured to selectively and controllably cause the ultrasonic transducers 107 in the array to transmit ultrasonic waves and to define / control parameters of the transmitted ultrasonic waves. The controller 101 is also configured to receive output data from other ultrasonic transducers in the array. In this manner, the ultrasonic transducers 107 are operated by the controller 101 to transmit and receive ultrasonic waves. In some implementations, the controller 101 is configured to electronically communicate directly with each ultrasonic transducer 107, while in other implementations, the controller 101 is indirectly connected to the multiple ultrasonic transducers 107 via a data collection and / or signal routing device (not shown) that is either incorporated into the controller 101 or provided as a separate add-on device.

[0029]

[0069] FIG. 2 illustrates a first example of a head-mounted transducer array device 201. The head-mounted device 201 includes a 3D-printed frame 203 sized to fit around the top of a subject's head. In this example, the 3D-printed frame 203 is generally tubular and supports a number of ultrasound transducers 205 arranged in an array around the inner circumference of the 3D-printed frame 203 such that when worn, the ultrasound transducer array surrounds the subject's head. Each ultrasound transducer 205 is communicatively connected to the controller 101 (of FIG. 1 ) via a cable 207 attached to the rear of the ultrasound transducer 205. In some implementations, the 3D-printed frame 203 can be custom sized according to the anatomical dimensions of the target subject. Additionally or alternatively, the position and orientation of the ultrasound transducers 205 supported by the 3D-printed frame 203 can be adjusted to target one or more specific locations within the brain of the target subject. In some implementations, this can be done by manually adjusting the position of the ultrasound transducers 205 to adjust the relative angle of the ultrasound transducers 205, while in other implementations, the 3D print frame 203 is configured to receive and support each ultrasound transducer 205 at a position and orientation specific to one or more target locations in the target subject's brain. Also, in some implementations, the relative angle of each transducer in the transducer array is known so that different target locations can be stimulated by selected different combinations of ultrasound transducers that intersect at the target locations (as described below). Thus, in some implementations, nearly any brain region can be targeted by selecting different combinations of ultrasound transducers to transmit ultrasound waves to brain tissue.

[0030]

[0070] FIG. 3 illustrates an alternative example of a head-mounted transducer array device 301. In this example, the head-mounted device comprises a frame 303 sized to be mounted around a subject's head 305, with two separate transducer arrays 307a, 307b disposed on either side of the 3D printed frame. As shown in FIG. 3, the ultrasound transducer arrays 307a, 307b are disposed on either side of the subject's head 305 when the head-mounted device 301 is worn by the subject. In the example of FIG. 3, the head-mounted device 301 also comprises an adjustable support mechanism 309 configured to allow adjustment of the positioning of the forehead support pad 310 and the nose support pad 311. The forehead support pad 310 and the nose support pad 311 can be adjusted to fit the anatomical shape of the target subject's head 305 to hold the head-mounted device 301 in place during use. In some implementations, the head mounted transducer array device (e.g., head mounted device 201 or head mounted device 301) is constructed from MRI compatible materials and cables, and use of the head mounted transducer array during MRI produces no detectable image distortion in the captured MR images, as shown in Figure 4. Additionally, in some implementations, the head mounted transducer array device (e.g., head mounted device 201 or head mounted device 301) may be waterproof.

[0031]

[0071] In one configuration, the device 301 comprises two spherical phased array transducers mounted in a plastic MRI compatible frame, positioned opposite each other and separated by a distance of 187 mm. The array elements are made of PMN-PT material with a surface area of ​​6 mm x 6 mm and operate at a fundamental frequency of 650 kHz. The two spherically focused arrays have a radius of 165 mm, 126 elements in a 9 x 14 element grid with 0.5 mm element spacing. Each array has a height of 55 mm and a width of 86 mm, and is 47.3 cm long. 2The transducers span an area of ​​100 mm. These transducers are configured to deliver ultrasound waves through the parietal and temporal bones of the subject. Specifically, the transducers are oriented parallel to the left and right sides of the subject's head. The transducers are driven by a programmable system (e.g., Vantage 256, Verasonics). The transducers are coupled to the subject with a hydrogel. Standard ultrasound coupling gel can be applied to the interface between the transducer and the hydrogel, and between the hydrogel and the head. The application of ultrasound gel is not critical given the presence of the hydrogel, but it can improve the transmission by approximately a factor of 2.

[0032]

[0072] In some embodiments, the system described herein is used to diagnose and treat neural sources of neurological or psychiatric disorders in a systematic and personalized manner. The head-mounted transducer array device 201, 301 operates to non-invasively modulate a designated brain target with high spatiotemporal resolution and in a multifocal manner. Furthermore, in some embodiments, the system is configured to compensate the skull, thereby depositing a deterministic amount of ultrasound energy in the designated brain target for effective and safe application. In some embodiments, the system is configured to apply ultrasound energy to the brain target using ultrasound superposition, as shown in the example of FIG. 5A using the head-mounted device 301 of FIG. 3.

[0033]

[0073] In some embodiments, the systems described herein are used to activate or release biocompatible nanoparticles carrying one or more therapeutic agents (e.g., propofol) for the treatment of a neurological or psychiatric disorder in a patient. The head-mounted transducer array device 201, 301 is operated to non-invasively modulate designated locations on the skull such that ultrasound can activate and release therapeutic agents at designated brain locations in the patient. Additionally, in some embodiments, the system is configured to compensate the skull (e.g., by utilizing the RTT method described herein) to provide a therapeutically effective release to the patient when ultrasound is applied to the skull.

[0034]

[0074] In the example of FIG. 5A, the system is operated to transmit ultrasound waves from three different ultrasound transducers of the first ultrasound transducer array 307a. The ultrasound waves are emitted at different relative trajectory angles such that the first ultrasound 501a, the second ultrasound 501b, and the third ultrasound 501c all intersect at the target location 503. At the target location 503, the superposition of the three ultrasound waves 501a, 501b, 501c applies targeted ultrasound energy to the target location 503. After passing the target location 503, the three ultrasound waves 501a, 501b, 501c continue along their respective trajectories until they reach the second ultrasound transducer array 307b. The output of the transducers of the second ultrasound transducer array 307b is monitored to measure the delivery strength of the ultrasound waves 501a, 501b, 501c that reach the second ultrasound transducer array 307b.

[0035]

[0075] FIG. 5A illustrates the operation of the head-mounted device 301 in a "free-field" region (i.e., where there is only water located between the ultrasonic transducer arrays 307a, 307b). As shown in FIG. 5B, when the head-mounted device 301 is placed on the head 305, the ultrasonic waves 501a, 501b, 501c converge to a location in the brain (e.g., target location 503) to apply ultrasonic energy through superposition. However, obstacles in the ultrasonic path between the first ultrasonic transducer array 307a and the second ultrasonic transducer array 307b affect the ultrasonic waves 501a, 501b, 501c, which in turn affect the ultrasonic energy applied at the target location 503. Obstacles in the ultrasonic path during use of the head-mounted device 301 can include, for example, anatomical features such as the subject's skull and scalp, hair, the coupling interface between the head-mounted device 301 and the head 305, and air bubbles in the coupling interface.

[0036]

[0076] For example, as shown in FIG. 5C, the ultrasound waves projected by the first ultrasound transducer array 307a must pass through a portion 509 of the skull before entering the brain and reaching a target location 503 within the brain. FIG. 6 shows an example of the relative intensity of ultrasound energy of each of the three ultrasound waves as they pass through the skull at 509. FIG. 7 shows the difference in the "actual intensity" at the target location 503 (i.e., the intensity of ultrasound energy at the target location 503 when the head-mounted device 301 is worn during use) compared to the "target intensity" at the target location 503 (i.e., the intensity of ultrasound energy at the target location 503 in the free-field region of FIG. 5A). FIG. 8 shows the ultrasound energy of one of the ultrasound waves projected from the first transducer j of the first ultrasound transducer array 307a as measured by the second transducer i of the second ultrasound transducer array 307b when the head-mounted device 301 is worn during use (as shown in FIG. 5B) and in the free-field region (as shown in FIG. 5A). These graphs demonstrate that the presence of the skull and other obstacles along the ultrasound path causes attenuation and phase dispersion (e.g., speedup) of ultrasound. Indeed, in some embodiments, the attenuation caused by the human skull prevents ultrasound from causing significant neural stimulation, an example of which can be observed in the red bar of FIG. 12 where the skull compensation proposed herein is not applied. Neural stimulation is restored when the compensation proposed herein is applied.

[0037]

[0077] To achieve a desired amount of ultrasound energy at a target location in the brain, the electronic controller 101 in some embodiments is configured to determine an appropriate compensation for the effects of the skull and other obstacles in the ultrasound path. In some such embodiments, the system uses ultrasound waves of the same frequency as those used for treatment to directly measure and compensate for ultrasound attenuation and phase dispersion due to all obstacles in the ultrasound path. In some embodiments, the required compensation value is established relativistically by contrasting ultrasound arrival times and amplitudes with the head present (FIG. 5B) and without the head (FIG. 5A). Using this mechanism to compensate for all obstacles in the ultrasound path (including, for example, the coupling interface of the head-worn device with the head, air bubbles, head anatomy, skull, brain tissue, etc.), the system can deliver ultrasound energy to the target location with the same intensity as that measured in water (without the head).

[0038]

[0078] FIG. 9 illustrates an example of a method executed by the controller 101 to determine the transmitted ultrasound beam and apply appropriate compensation to achieve a desired / target ultrasound energy at a target location in the brain. First, ultrasound energy is measured via a "free-field" sequence (step 901) in which ultrasound is transmitted by a first set of transducers such that the ultrasound passes through a "free-field" region (e.g., only water between the ultrasound transducers) and is measured by a second set of transducers on the opposite side of the ultrasound path. Second, the head-mounted device 301 is placed on the head and a "through" sequence is measured (step 903) in which the same ultrasound is transmitted from the same first set of transducers and measured by the same second set of transducers; however, in the "through" sequence, the ultrasound path travels through the subject's head / skull. The measurements from these two sequences are compared to quantify the difference in attenuation and phase values ​​(step 905). Appropriate phase and amplitude adjustments are determined based on the difference in attenuation and phase values ​​(step 907). The ultrasonic waveform to each transducer is applied with the determined amplitude and phase adjustments, step 909. Further details regarding the "transmission" methodology are provided in Example 2 below.

[0039]

[0079] For measurements in water (e.g., free-field sequence measurements), an obstacle in the ultrasound path (e.g., the skull) reduces the beam emitted from each transducer element i by a factor A i and time the beam with a relative time τ i The compensation method of FIG. 9 estimates and corrects these values, and then increases the amplitude of each ultrasonic beam by a factor of 1.

[0040]

number

[0041] Scale the emission by τ i In some embodiments, the speedup time τ i and attenuation A i The reconstruction of is solved separately.

[0080] For damping, in some implementations, the controller 101 is configured to solve the following system of equations:

[0042]

number

[0043] In the formula, A ij is the relative attenuation measured by ultrasound transmission through both sides of the skull.

[0044]

number

[0045] represents the extended path that the ultrasound travels through the skull with an angle β between transducer elements i and j. Because the attenuation through two opposing segments of the skull doubles, a logarithmic formula is used for the attenuation. This system of linear equations can be expressed in matrix form as Kx=b, where K is the sum of k ij matrix of coefficients, x is the desired value x=[A1,A2,...,A 256 ], and b is the vector of measurements A ij The solution x minimizes the sum of squared errors (b-Kx)'×(b-Kx).

[0046]

[0081] In some embodiments, the controller 101 may control a phase shift as follows:

[0047]

number

[0048] For every transducer pair, h ij (t) corresponds to the received signal of the i-th transducer after a short pulse is emitted from the j-th transducer.

[0049]

number

[0050] (i.e., free-field measurements) and through-the-skull measurements

[0051]

number

[0052] For each receiving transducer i of the total N transducers, the controller 101 calculates the transmission delay in water (i.e., free-field measurement)

[0053]

number

[0054] This determines the total transmit and receive waveform.

[0055]

number

[0056] is focused on element i.

[0082] Timeshift

[0057]

number

[0058] The sum of these waveforms in any vector in is expressed as follows:

[0059]

number

[0060] During the ceremony,

[0061]

number

[0062]

[0083] The purpose is to delay the speed of water

[0063]

number

[0064] Each wave received at element i in the water is delayed to compensate for its speed-up in front of the transmitting element.

[0065]

number

[0066] compared to the waves passing through the skull after application of

[0067]

number

[0068] In some implementations, the controller 101 calculates these delays by optimizing the following equation:

[0069]

number

[0070] The method is configured to identify:

[0071]

number

[0072]

[0084] FIG. 10 shows the intensities obtained with the compensation method of FIG. 9 at three different target locations in the brain: the superior lateral branch of the medial forebrain bundle, the ventral internal capsule / ventral striatum, and the ventral intermediate nucleus (VIM). For each experiment, the graphs show the intensity of ultrasound energy at the target location without the skull (i.e., free-field measurements), the intensity at the same target location when the same ultrasound beam is applied to the skull (i.e., transmission measurements), and the intensity at the same target location when a compensated ultrasound beam is applied to the head (i.e., "diadem-corrected" stimulation measurements). The graphs represent a quantification (mean ± s.e.m.) of the compensation accuracy in eight different samples. As shown in these graphs, the target peak intensity (14 W / cm2) delivered to each target was 1.25 W / cm2. 2 ) is severely attenuated by the skull for all targets (see the red bars in Fig. 10), but the compensation method in Fig. 9 can accurately recover the target intensity values ​​(see the green bars in Fig. 10).

[0073]

[0085] 11 and 12 show the results of another set of experiments using the compensation method of FIG. 9. The head-mounted device 301 was configured to target the nerves and nerve endings of the thumbs of 11 participants. The thumbs were fixed at the location of the VIM in an ex-vivo skull. The device delivered 300 ms stimulation to the targets at a frequency of 650 kHz. Data was collected without the skull, through the ex-vivo skull, and through the skull after applying skull compensation. The experiment also included applying stimulation 10 mm below the finger (i.e., "off-target" sham stimulation). Subjects closed their eyes and wore noise-canceling headphones. Subjects were blinded as to whether compensation, uncompensation, or sham stimulation was applied, and reported any nociceptive responses. A nociceptive response is indicative of stimulation of the nerves and nerve endings.

[0074]

[0086] The graph in FIG. 11 shows the dose-response relationship of the stimulation. A two-way analysis of variance (ANOVA) detected significant modulation by ultrasound pressure (F(2,60)=25.11, p<0.001). The skull compensation was accurate as there was no significant difference between the responses to skull-free and skull-compensated stimulation (p=0.96). The graph in FIG. 12 shows the quantification of the effect at the highest applied stimulation pressure (1.33 MPa). The p-values ​​indicate the significance of the respective two-tailed t-test. As shown in FIG. 12, the response rates were nearly identical for skull-free and skull-compensated stimulation, whereas the response rate for uncompensated stimulation applied through the skull was nearly zero. This result clearly demonstrates the importance of compensation: without compensation, there is essentially no neural stimulation. Furthermore, the response rate for off-target stimulation was also nearly zero, suggesting that the stimulation applied by the head-mounted device 201, 301 likely does not have a discernible stimulation effect outside of the targeted target area.

[0075]

[0087] Figures 13 and 14 show the results of yet another stimulation experiment using a head-mounted transducer array device. Figure 13 shows the measured peak pressures at each of three different target sites (ventral tegmental area (VTA), ventral internal capsule / ventral striatum (VC / VS), superior lateral branch of medial forebrain bundle (slMFB)) for free-field measurements (i.e., "target (without skull)"), uncompensated target transmission measurements (i.e., "actual (with skull)"), and compensated target transmission measurements (i.e., "compensated (US corrected)"). This figure confirms the concept of Figures 10-12 that the compensation procedure described in Figure 9 is critical for deterministic and effective delivery of ultrasound. Figure 14 shows the quantification of focal volumes for three identical target regions under three identical stimulation conditions. As illustrated by Figure 14, the presence of the skull in the ultrasound path not only attenuates the ultrasound but also increases the focal volume of the stimulation imparted by the superimposed ultrasound beams. However, FIG. 14 also shows that the focal volume can be reduced by applying compensation according to the method of FIG. 9. Thus, FIG. 14 shows that in addition to compensating for attenuation and phase shift caused by obstacles in the ultrasound path, the method described above with reference to the example of FIG. 9 can also be configured to adjust and fine-tune the focal volume of the applied stimulation. In some embodiments, the controller 101 is configured to collaboratively optimize the adjusted compensation applied to the transmitted ultrasound beam to bring both the corrected peak pressure and the focal volume as close as possible to the target values. Under some conditions, adjustments that bring the focal volume closer to the target can cause the corrected peak pressure to move further away from the peak pressure target, and if both targets cannot be achieved simultaneously, a variety of different optimization techniques can be employed by the controller 101 to determine an optimized solution to balance the deviations of the peak pressure and focal volume from the targets.

[0076]

[0088] FIG. 15 shows an alternative example of a head-mounted transducer array device 301. In this example, the head-mounted device comprises a frame sized to be worn around the subject's head, with two separate transducer arrays (black) positioned on either side of the frame. As shown in FIG. 3, the ultrasound transducer arrays are positioned on either side of the subject's head when the head-mounted device is worn by the subject. In the example of FIG. 3, the head-mounted device 301 also comprises an adjustable support mechanism configured for adjustable positioning. In some implementations, the head-mounted transducer array device (e.g., head-mounted device 201 or head-mounted device 301) is constructed from MRI compatible materials and cables, and as shown in FIG. 17C-17G and FIG. 33, the use of the head-mounted transducer array during MRI does not cause detectable image distortion in the captured MR images. Additionally, in some implementations, the head-mounted transducer array device may be waterproof. EXAMPLES

[0077]

[0089] Demonstration of effective neuromodulation in humans using a head-mounted device

[0090] The device was applied to patients with treatment-resistant depression to modulate the subcingulate cortex, a deep brain structure. The effect on the target was verified using fMRI BOLD. Furthermore, target modulation improved the patients' mood state. This effect was specific to the stimulated target and was not observed with sham stimulation, which delivered the same pressure and waveform of stimulation to the brain but was not focused.

[0078]

[0091] A head-mounted device 301, shown in Figure 5B, was applied to a patient (see Figure 15). The device contained two sets of 126-element phased array transducers placed on either side of the patient's head. The transmit and receive capabilities of each transducer array element allow for ultrasound-assisted correction for the attenuation (and phase dispersion) of each ultrasound beam. A single pre-operative MRI and mechanical resist were used to guide the transcranial application of focused ultrasound to multiple regions of the patient's brain. When focused on a target in the infracingulate cortex, the array measured 20.4 mm axially, 2.4 mm lateral, and 3.6 mm in elevation, for a total volume of 142.71 mm. 3 (corresponding to the volume of a sphere with a diameter of 6.48 mm) (see Figure 16).

[0079]

[0092] Figures 17A-B show the intensity field generated by the array superimposed on the subject's brain anatomy for scale. The phased array geometry allows for flexible electronic focus steering with dimensions of ±45mm in the axial dimension, ±25mm in the lateral dimension, and ±15mm in the elevation dimension (Figure 23). An MRI-compatible plastic frame (shown in Figure 15) holds the array, allowing it to slide on horizontal and vertical tracks and then lock into position anywhere on the side of the subject's head after the array has been moved.

[0080]

[0093] After correction with ultrasound to resist and skull, the subgenual cingulate cortex of the brain was sonicated while measuring fMRI BOLD response as in Figures 17C-17G. Figures 17C and 17D show significant clear activation in the target region (peak level: p=0.003, t=5.52, ZE=5.37, cluster level p<0.001 (family weighted error corrected, kE=69 voxels)). This fMRI BOLD activity was specific to the onset of ultrasound stimulation (Figure 17E). Contrast estimate of BOLD activity in this region from ultrasound off to ultrasound on condition shows a large effect size of 3.07±0.56 (mean±sem). Subjects blinded to type of stimulation were also presented with active sham stimulation of unfocused plane wave ultrasound as a negative control. This unfocused beam had the same stimulation parameters and intensity but phase delay was adjusted to focus the energy far outside the subject's head. Under this active sham stimulation condition, no significant BOLD activity was measured near the target (Figures 17F-G).

[0081]

[0094] Outside the MRI scanner, subjects were presented with 10 sonications at 1 MPa pressure, 30 ms on pulse time, and 4 s pulse intervals, varying stimulation duration and focal position. Improvements in subjects' self-reported ratings of depression, anxiety, and emotional valence were observed when the subgenual cingulate cortex was sonicated for 150 to 300 s (Figure 18). Subjects reported no change in mood when sonicated to the rostral ventral striatum or with unfocused plane wave stimulation of the same intensity. This effect was also specific to stimulation duration, with changes in mood scores increasing with increasing stimulation duration. Stimulations of less than 150 s resulted in little change in mood scores. Aside from these effects on mood, subjects reported the subjective experience of being able to stay in train of thought for longer during stimulation and feeling hopeful about future events for the first time in two weeks. Importantly, subjects reported no adverse effects during the three 90-min sonication sessions (Table 1).

[0082] [Table 1]

[0083]

[0095] Prior to sonicating a particular target, transmission measurements through the subject were obtained, the attenuation and phase delay in front of each element were estimated compared to water, and stimulation parameters were adjusted to compensate for these ultrasonic aberrations. Figures 5A-5C, 6, 7, 8A-8B, and 19 show the steps involved in the compensation and examples of transmission waveforms acquired in water and through the subject. Taking a full transmission scan across all 252 elements takes less than 1 second. Importantly, this scan measures attenuation from all sources of attenuation in the beam path, namely coupling, hair, local air pockets, skull, and brain. An instantaneous measurement of pressure transmission through the skull for each element informed the operator of the quality of acoustic coupling and allowed uncoupled elements to be fixed before sonication. Figure 20 shows the average amplitude scaling factor applied to the transducer elements to compensate for the estimated attenuation. The average amplitude scaling factor was 4.84 ± 0.94 (mean ± SD) for stimulation session 1, 5.59 ± 1.09 for session 2, and 5.68 ± 0.97 for session 3. All elements with an estimated pressure transmissibility below 12% were switched off and the remaining elements had their amplitudes increased to compensate for their lost contribution.

[0084]

[0096] Finally, to assess the repeatability of targeting, the targeting error of the device was measured across multiple sessions and multiple subjects. The plastic frame allowed repeatable positioning of the ultrasound transducer. Across all sessions and subjects, the ultrasound transducer position varied 0.89 ± 0.64 (mean ± SD) overall, 0.45 ± 0.32 in the x dimension, 0.43 ± 0.14 in the y dimension, and 0.44 ± 0.17 in the z dimension. The thermoplastic mask ensured that the subject's head was fixed in the same position across multiple sessions. The subject's fiducial marker position varied 1.28 ± 0.66 overall, 0.53 ± 0.19 in the x dimension, 0.68 ± 0.27 in the y dimension, and 0.71 ± 0.31 in the z dimension, averaged across multiple trials for all subjects (Figure 21). The targeting error, i.e., the deviation between the location of the virtual ultrasound focus and the subject's brain target, was assessed by measuring the average difference in location between a fiducial point on the subject and a fiducial point on the ultrasound transducer. Overall, the average targeting error of the device was 1.64 ± 0.66 across subjects, 0.77 ± 0.50 in the x dimension, 0.93 ± 0.41 in the y dimension, and 0.99 ± 0.49 in the z dimension (Figure 22).

[0085]

[0097] In this study described in this example, the effectiveness of the device was verified by producing significant changes in fMRI BOLD activity in the target region. The device compensates for attenuation and phase shifts due to the subject's skull, hair, and acoustic coupling; thereby improving current ultrasound neuromodulation techniques by delivering effective and safe intensity to the target. Figure 33 (left) shows that this feature is important for effective ultrasound neuromodulation. In the non-application case (Figure 33, right), which was the case when using the existing ultrasound device, there was no neuromodulation effect.

[0086]

[0098] The mechanical frame and phased array system allow for flexible focusing of energy to spatially specific targets deep in the brain. With a novel mechanical resist approach, the device enables reproducible targeting of deep brain regions across subjects and treatment sessions, outside of an MRI scanner, without the need for expensive neuronavigation systems.

[0087]

[0099] The device provides effective neuromodulation in humans that can be monitored by fMRI BOLD measurements. A clear fMRI bold response was elicited in the target when stimulating the infracingulate cortex of human subjects. Changes in fMRI BOLD activity were temporally synchronized to ultrasound stimulation and were not seen during sham stimulation. Taken together, these data provide strong evidence that the device locally activates targets deep in the brain and that the response is specific to ultrasound stimulation. The fMRI BOLD readout available in this MRI-compatible device gives valuable feedback on the amplitude, polarity, and targeting precision of the neuromodulation. The stimulation parameters of the MRI scanner ensure that stimulation is delivered to the target and found to suppress activity in the infracingulate cortex.

[0088]

[0100] Modulation of the infragenual cingulate cortex effectively induced positive changes in subjects' mood states. Specifically, improvements were observed in subjects' self-reported depression, anxiety, and affective valence scores. Subjects reported no changes in mood to sham stimulation, stimulation of the rostral ventral striatum, or the majority of stimulations lasting less than 1 minute. Overall, mood changes were specific to both the target region (infragenual cingulate cortex) and stimulation parameters (durations greater than 1 minute). Stimulation with the device was safe and well tolerated. Subjects reported no adverse effects over three 90-minute stimulation sessions.

[0089]

[0101] The ability of the device to measure the attenuation of the ultrasound beam is important for both safety and efficacy (Figure 33). Ultrasound attenuation varies from session to session. By repeatedly placing the device on the same section of the skull, these differences are likely due to variability in the coupling of the transducer to the subject's head. A safety feature of the device is to turn off elements with particularly high attenuation, which could be due to local air pockets or particularly thick areas of the skull. In either case, sonication can result in dangerous cavitation or heating effects. With regard to efficacy, the considerable amplitude magnification and phase compensation applied to the array elements was important for stimulation. In titrating doses, several ultrasound waves were delivered through the skull with a delivery pressure of 0.4 MPa and no change in mood scores was observed. Considering that the mean pressure amplitude magnification for this particular subject was 5.37, it is likely that a target intensity of 33.78 W / cm would have been achieved if no skull compensation had been performed. 2 (1MPa), 1.09W / cm 2 (0.18 MPa). CT correction of the skull, which is effective for measuring phase dispersion caused by the skull but not for measuring attenuation, cannot account for attenuation due to hair, entrapped air bubbles, and variations in acoustic coupling. These additional barriers are substantial, attenuating the transmitted intensity by approximately 0-64% due to approximately 20% due to hair and up to 100% coupling due to localized air bubbles or air pockets. For example, in some sessions, as many as 60 of 252 elements were found to be not acoustically coupled to the subject prior to sonication, and this coupling problem had to be corrected prior to sonication. The remaining uncoupled elements were turned off. This feedback is not present in the CT correction, leading to potentially dangerous insonation of localized air pockets and less than effective pressure at the target.

[0090]

[0102] The flexible and reproducible targeting of the device allows for intervention in many regions throughout the brain (Figure 23). To affect the infracing cingulate cortex of this subject, the beam was electronically steered 17 mm laterally and 9 mm in height from the geometric center of the array. By locking the array in place and using electronic beamforming, superior and inferior sections of the infracing cingulate cortex, the ventral striatum were insonified, and the beam was steered outside the head for sham stimulation, all during the same stimulation session. With the ability to beamform to different targets in microseconds, hundreds of unique brain regions can be stimulated by the phased array system every minute. Other regions of the brain such as the thalamic nuclei, amygdala, cingulate cortex, insular cortex, nucleus accumbens, and ventral tegmental area can be addressed through a combination of physical movement of the array and beamforming of the focus to a given target. This fast and flexible targeting is unique to phased array systems. Neuromodulation devices have typically used single-element transducers under MRI guidance for targeting, or optical neuronavigation. The mechanical resist method allows precise targeting of deep brain regions both inside and outside the MRI scanner, without the need for such expensive resist tools.

[0091]

[0103] Although fMRI BOLD activity to stimulation has been robustly demonstrated in animal studies, only two groups have reported fMRI measurements in humans. We extend these results by showing a relatively strong effect size, along with the first demonstration of fMRI BOLD response to ultrasound stimulation in a relatively deep region, the subgenual cingulate cortex. The strong fMRI activation is due to the increased target pressure obtained by correcting for cranial aberrations (Figure 33). Overall, these significant and temporally synchronized BOLD responses to ultrasound stimulation indicate the important capabilities of the device to monitor the magnitude, polarity, and targeting precision of neuromodulation effects.

[0092]

[0104] This study is also the first to demonstrate that tFUS stimulation of the infracingulate cortex improves mood state in patients with treatment-resistant depression. Previous studies have similarly shown improvement in mood with FUS stimulation of the ventral lateral prefrontal cortex and inferior frontal gyrus. Not only was an immediate improvement in mood observed, but subjective effects such as hope for future events were also observed. The rationale for targeting the infracingulate cortex comes from previous DBS studies, as well as from the neuroscience literature implicating this region in major depressive disorder. The results of this study support the notion that the SGC controls mood state and is a promising target for more prolonged FUS stimulation.

[0093]

[0105] The fMRI and mood response results of this study were limited to one subject in an ongoing clinical trial with a total of 20 subjects. However, the fMRI results were statistically significant from baseline, and the mood effects were reproducible across sessions and robust to sham stimulation; therefore, these data demonstrate proof of concept in the first human subject of this study. Stimulation targets in this first subject were limited to the ventral striatum and subgenual cingulate cortex. Due to the fast and flexible beamforming of ultrasound, multiple targets within the unique brain targets of the medial forebrain bundle, ventral tegmental area, and anterior cingulate cortex can be sonicated in rapid sequences at nearly the same time. The device is limited to sonication of the penetration area where a cranial penetration pathway can be established, i.e., both the left and right sides of the head. Thus, although the device can access nearly the entire subcortical volume of the brain through mechanical movement of the array and phased array manipulation, access to brain cortical targets is limited.

[0094]

[0106] A non-invasive device is described herein that can controllably deliver ultrasound for safe and effective deep brain stimulation. The combination of ultrasound stimulation with fMRI allows for easy monitoring of neuromodulation. With regard to practicality, a single T1 MRI of the patient's head in the device is required for precise targeting in all subsequent sonication procedures. Furthermore, no shaving of the head is required. The phased array system can adjust the stimulation location in the brain in microseconds with millimeter accuracy, stimulating hundreds of unique brain targets per minute. Compared to existing FUS brain stimulation devices (BXPulsar, NeuroFUS) or surgical (Exablate Neuro) devices, this is the only device that has the capability to compensate for the variability of the human skull, hair attenuation, and acoustic coupling. This feature is important because each barrier distorts and attenuates ultrasound waves intensely and unpredictably. Future applications of low-intensity ultrasound to the brain must address this challenge so that the ultrasound energy delivered to the target is safe, effective, and reproducible from patient to patient. The precise correction of these distortions by the present device is expected to significantly improve the safety and efficacy of not only neuromodulation but also other low-intensity applications of transcranial focused ultrasound, such as localized drug delivery and blood-brain barrier opening, which, like ultrasound neuromodulation, are highly dependent on the delivered ultrasound intensity. EXAMPLES

[0095]

[0107] Testing the accuracy and validity of RTT

[0108] Figure 24A shows the severity of acoustic attenuation by the skull. In 8 ex vivo skulls, ultrasound intensity delivered to deep brain locations was found to be attenuated by a factor of 11.4 ± 6.8 (mean ± SD), replicating previous findings. In principle, attenuation could be estimated using tabulated values ​​(e.g., 20, 21), but such estimates are imprecise and uncertain due to the large individual variability in attenuation (Figure 24B). For example, using the values ​​from these two studies would lead to an overestimation of the mean value by a factor of 1.7 ± 0.67 and an underestimation of the mean value by a factor of 0.63 ± 0.25, respectively, leading to a large variability (pooled standard deviation equal to 0.47 for a normalized intensity of 1.0). Compensation of ultrasound phase dispersion, which can be obtained using existing methods, is useful for ultrasound-based surgery and to some extent for the current purpose of delivering deterministic intensities to specific targets for repeated applications (Figure 24C). Nevertheless, even ideal correction of phase assumptions based on ground truth measurements (Figure 24C) leaves an average 85% discrepancy between the desired intensity and the actual intensity delivered to the brain target.

[0096]

[0109] The system of FIG. 1A and the device shown in FIG. 5A were utilized to test the accuracy of the RTT described above when ultrasound was focused on a specific target in a human ex vivo skull. The induced field using hydrophones was measured. The measured intensity in four conditions was evaluated. First, the intensity in a free field was measured, which is free of obstacles to the ultrasound. This intensity corresponds to the target intensity delivered by the operator to the target. Second, the skull was placed between the device and the hydrophones and the resulting intensity was measured. In this case, it represented the worst case scenario without any correction for the skull. Third, a hypothetical ideal correction for the skull was evaluated. To do so, a hydrophone was used to measure the attenuation and phase dispersion of each element of the device, thereby obtaining ground truth values. These ground truth values ​​were used to compensate for these aberrations while scaling the magnitude of the ultrasound emitted from each element and delaying it accordingly, as if the skull was not present. And fourth, the RTT correction was applied.

[0097]

[0110] These measurements were made inside a human ex vivo skull that had been degassed by immersion in water for 8 h. Figure 25 shows the spatial peak intensity and associated field of a target placed at the center of the two transducers. This figure reinforces the notion that the human skull significantly attenuates the intensity (red) delivered to the brain. Compared to the free-field values, the ultrasound intensity transmitted through the skull was attenuated by a factor of 11.4 ± 6.8 (mean ± SD), down to 10.7 ± 4.2% of the target intensity. The difference between the free-field and skull transmission values ​​was significant (t7 = 60.7, p = 8.6 × 10-11, paired two-tailed t-test).

[0098]

[0111] RTT with phased array was then applied. Figure 25 shows that RTT recovered the target intensity value (green). The RTT-corrected intensity was 98.8 ± 17.8% (mean ± SD) of the target value in the free field, with no significant difference between the means of the two conditions (t7 = 0.18, p = 0.86, paired two-tailed t-test). The mean was also not significantly different from the hypothesized best-corrected value based on ground truth measurements of the hydrophone inside the skull (black bars; t7 = 0.17, p = 0.86, paired two-tailed t-test). One skull (purple data points) significantly attenuated the ultrasound (attenuation coefficient 26.9). This was likely due to a visually present protrusion associated with hyperostosis, as assessed by the neurosurgeon. For this skull, the RTT correction was less accurate, resulting in 0.62 times the target intensity.

[0099]

[0112] Next evaluated was the applicability of RTT to the human head, which represents even more significant barriers for transcranial ultrasound, including hair, scalp, acoustic coupling, and the brain. The study also evaluated the safety of the method. RTT was designed to be safe. RTT scans consist of short (<100 μs) low intensity (mean peak pressure 80 kPa in free field) ultrasound pulses. RTT scans took less than 1 second to complete. Subjects (n=5) experienced no discomfort during the procedure. Figure 26, blue, shows the average transmission attenuation through both sides of the head, separately for each subject. This figure demonstrates that the method provides transmission quality comparable to the ex vivo skull (gray). Specifically, the receive wave elements on both sides of the head recorded an average of 7.3 ± 4.8% (mean ± SD, n = 5 subjects) of the signal amplitude when the RTT was applied through the human skull and 12.6 ± 8.2% (mean ± SD, n = 8 skulls) in the characterized ex vivo human skull. This additional attenuation coefficient of 1.7 is expected because the application of ultrasound through the human head incurs additional attenuation due to the scalp, hair, coupling, air bubbles or air pockets, as well as intracranial tissues. At 1 week follow-up, no subjects reported any side effects. Thus, the RTT can be safely applied to the human head and allows for direct measurement of attenuation due to all obstacles in the ultrasound path.

[0100]

[0113] We next tested the robustness of RTT with respect to brain target location by using a phased array to refocus ultrasound on the target, covering the full operating range of the device, i.e., 10 mm axial, 20 mm axial, 10 mm lateral, 20 mm lateral, and 15 mm elevation relative to the central target (Figure 27). With RTT correction, the delivered intensity was 96.3 ± 21.4%, 94.8 ± 23.2%, 92.8 ± 16.4%, 62.5 ± 15.7%, and 71.6 ± 18.03% of the target value, respectively, for each target. There were no significant differences between the mean target peak intensity and the RTT-corrected peak intensity for the central target (t7 = 0.18, p = 0.86, paired two-tailed t-test), axial 10 mm (t7 = 0.48, p = 0.64, paired two-tailed t-test), lateral 10 mm (t7 = 0.42, p = 0.55, paired two-tailed t-test), and axial 20 mm (t7 = 1.23, p = 0.26, paired two-tailed t-test). There were significant differences in the mean delivered intensity for targets at lateral 20 mm (t7 = 6.748, p = 0.0002, paired two-tailed t-test) and elevation 15 mm (t7 = 4.5, p = 0.003, paired two-tailed t-test).

[0101]

[0114] The relative contributions of attenuation and phase dispersion, two important components of ultrasound aberration due to the skull, were then examined. Figure 28 shows the spatial peak intensity for each correction type alone and in combination. For the central target, phase-only correction resulted in a mean intensity of 13.8 ± 4.3% (mean ± SD) for ideal hydrophone correction (gray) and 11.9 ± 4.9% for RTT (green). No correction (red) was 10.7 ± 4.2% of the free-field intensity. With amplitude-only correction, the spatial intensity peaks were 71 ± 12.5 for hydrophone and 93.8 ± 28.9 for RTT. Thus, correction for attenuation (i.e., for the amplitude of the received signal) constitutes an important factor in the delivered ultrasound intensity. The inclusion of phase correction is even more desirable in that the resulting combined correction (Figure 28(c)) results in a mean delivered intensity of 98.8 ± 17.8%.

[0102]

[0115] The robustness of RTT to specific hardware was further tested. In particular, RTT was performed on an array with the same number of elements but a much larger aperture (Figure 29). In this configuration, the skull (n=4 samples) reduced the intensity at the geometric center to 6.3±1.7% of the target free-field value, similar to Figure 24.

[0103]

[0116] RTT compensation restored the intensity at the target to 104 ± 18.1% of the target value. After compensation, there was no significant difference between the target intensity and the mean RTT-recovered intensity (t3 = 0.47, p = 0.67, paired two-tailed t-test). EXAMPLES

[0104]

[0117] Demonstration of effective application of transcranial ultrasound for neuromodulation using a head-mounted device in humans

[0118] To test the effect on nerves in intact biological tissue, 11 human subjects were instructed to place their thumbs on a central target holder in an in vivo skull. Subjects' responses to ultrasound with and without RTT applied were quantified (see Methods). The targets were stimulated for 300 ms with specific pressure levels, and the effect on the subjects' nociceptive response was assessed. A nociceptive response refers to the stimulation of nerves or nerve endings in tissue. RTT was found to be important for effective stimulation (Figure 30(a)). No significant stimulation was observed without RTT (red; t11=1.00, p=0.34, one-sample two-tailed t-test). After RTT, the subjects' response rate to stimulation reached 62.7%. This level was statistically equivalent to the 66.3% response rate obtained with the best ground truth correction of the assumptions (t10=0.58, p=0.57, paired two-tailed t-test), which was comparable to the case without the skull. FIG. 31 shows the individual responses to each correction for all subjects.

[0105]

[0119] To control for possible confounds that may be related to ultrasound stimulation, we randomly interleaved a sham stimulation in which ultrasound was delivered 10 mm below the target with hydrophone compensation. This off-target stimulation did not result in significant stimulation (yellow, p = 0.19, one-sample two-tailed t-test, t = 1.39). This controls for potential artifact effects and confirms the spatial specificity of the stimulation.

[0106]

[0120] The dose-dependence of the stimulation effect was further investigated. Specifically, stimulation across three intensity levels was varied. It was found that the stimulation effect increased as the level of ultrasound increased (Figure 30(b)). The response frequency reached 62.7% for the strongest stimulation (1.8 MPa) and was also significant for the weakest stimulation tested (1.3 MPa; t10=7.63, p=1.7×10-5, one-sample two-tailed t-test). The effect of stimulation level was highly significant (2-way ANOVA, F2,60=25.24, p=1.1×10-8). The responses were statistically indistinguishable from the hypothesized ideal correction (green vs. black; 2-way ANOVA, F1,60=0.41, p=0.52) and there was no significant interaction between the two factors (F2,60=0.20, p=0.98).

[0107]

[0121] Thus, precise compensation of delivery intensity to brain targets (Figures 25, 27) also leads to restoration of stimulation effects at response levels that would not be achievable without the RTT method described herein (Figure 30).

[0108]

[0122] The measurements and compensation for the human head described herein are important for effective ultrasound neuromodulation in the human brain. The left side of Figure 33 shows that compensation for the head enables strong ultrasound neuromodulation. The right side of Figure 33 shows that when this compensation is not applied, as in existing devices and approaches, there is no significant ultrasound neuromodulation. EXAMPLES

[0109]

[0123] Demonstration of effective application of transcranial ultrasound for drug delivery using a head-mounted device in humans

[0124] RTT was tested whether it could be used to release therapeutic agents (e.g., any hydrophobic drug such as propofol, mycophenolate mofetil, and ketamine) in clinically relevant and deterministic doses to specific locations within the skull. In one example, an ultrasound-sensitive nanoparticle carrier was devised and the neuromodulation drug propofol was encapsulated in nanoparticles at a concentration of 0.063 mg / ml. How the nanoparticles responded to ultrasound when RTT correction was applied and not applied was tested in a manner similar to Figure 30. It was found that RTT was important in mediating effective release when ultrasound was applied to the skull (Figure 32(a)). Without RTT (red), the amount of drug detected was not different from that without stimulation (purple) (t14=0.30, p=0.77, two-sample two-tailed t-test). Application of RTT (green) nearly tripled the release effect (2.9-fold increase), with 31.6% of the encapsulated propofol being released. This level was statistically equivalent to the 31.8% release obtained with the best correction of the assumptions (black) (t18=0.08, p=0.94, paired two-tailed t-test).

[0110]

[0125] The spatial specificity of release was confirmed using sham conditions where ultrasound was focused 10 mm below each vial (Figure 32(a)). In this case, the amount of drug detected did not differ from when ultrasound was not applied (purple) (t14=0.30, p=0.77).

[0111]

[0126] The dose dependence of the release was also investigated. To do so, the delivered ultrasound intensity was varied at the same levels as in Figure 30. An increasing stimulation effect was observed as the ultrasound level increased (Figure 32(b)). The effect of stimulation level was highly significant (2-way ANOVA, F2,54=84.53, p=2.3×10-17). The release level was statistically indistinguishable with the hypothetical ideal correction (green vs. black; 2-way ANOVA, F1,54=0.02, p=0.89), and there was no significant interaction between the two factors (F2,54=0.35, p=0.7).

[0112]

[0127] Thus, precise compensation of delivery intensity to the brain target (Figure 25, Figure 27) also leads to recovery of the stimulation effect at a level of response that would not be achievable without this method (Figure 30). In other words, the target ultrasound intensity is sufficient to activate the microbubbles, thus temporarily destroying the blood-brain barrier, for local delivery of drugs, genes and stem cells across the blood-brain barrier.

[0113]

[0128] Thus, in various different embodiments, the present disclosure provides, among other things, systems and methods for using ultrasound to apply focused ultrasound to locations in the brain, and using the ultrasound itself to precisely compensate for attenuation, phase shift, and / or focal volume changes due to the presence of a head in the ultrasound path. Additional features and advantages of the invention are set forth in the following claims and accompanying drawings.

Claims

1. 1. An ultrasound delivery system comprising: a head-mounted device including at least one array of ultrasound transducers including ultrasound transducers positioned on opposite sides of the volume; a controller, determining a first set of ultrasound waves to be transmitted by a first set of ultrasound transducers of the at least one array of ultrasound transducers to achieve a target ultrasound stimulation energy at a target location within the volume corresponding to a target brain location; Driving the first set of ultrasonic transducers according to the determined first set of ultrasonic waves; capturing free-field measurements of ultrasound using a second set of ultrasound transducers on an opposite side of the volume from the first set of ultrasound transducers while the head-worn device is not applied to the head; capturing ultrasound transmission measurements using the second set of ultrasound transducers while the head-worn device is applied to the head; comparing the free-field measurements with the transmission measurements to determine attenuation and phase shift of the ultrasound waves due, at least in part, to the presence of a skull in an ultrasound path between the first set of ultrasound transducers and the second set of ultrasound transducers; determining an adjusted set of ultrasound waves transmitted by the first set of ultrasound transducers that compensates for the determined attenuation and the determined phase shift to achieve the target ultrasound stimulation energy at the target location within the volume while the head-mounted device is applied to the head; driving the first set of ultrasonic transducers according to the determined adjusted set of ultrasonic waves while the head-mounted device is applied to the head; a controller configured to 1. An ultrasound delivery system comprising:

2. the at least one array of ultrasound transducers includes a plurality of ultrasound transducers positioned and oriented to each project an ultrasound beam along a different ultrasound path; the controller is further configured to identify the first set of ultrasound transducers by identifying ultrasound transducers having ultrasound paths that intersect at the target location; The controller is configured to determine a combination of ultrasound waves to be transmitted by the first set of ultrasound transducers to achieve the target ultrasound stimulation energy at the target location by determining a combination of ultrasound waves that combine by superposition when the ultrasound paths intersect at the target location to generate the target ultrasound stimulation energy. The ultrasound delivery system of claim 1 .

3. 2. The ultrasound delivery system of claim 1, wherein the controller is further configured to determine the attenuation and the phase shift due to ultrasound passing through the skull once based in part on a difference between the measured ultrasound passing through the free-field volume and the measured altered ultrasound passing through the head, the difference being indicative of the attenuation and phase shift caused by the ultrasound passing through each segment of the head.

4. the controller is further configured to determine a plurality of cranial attenuation values ​​including a determined cranial attenuation value for each ultrasound transducer of the first set of ultrasound transducers, each cranial attenuation value of the plurality of cranial attenuation values ​​representing attenuation of the ultrasound waves transmitted by a respective one of the ultrasound transducers through the skull; the controller is configured to determine the adjusted set of ultrasound waves by scaling the amplitude of the ultrasound waves transmitted by each ultrasound transducer of the first set of ultrasound transducers by the inverse of the determined cranial attenuation value corresponding to each of the ultrasound transducers. The ultrasound delivery system of claim 1 .

5. Determining the plurality of cranial attenuation values ​​comprises: [Equation 1] calculating the cranial attenuation value for each ultrasound transducer of the first set of ultrasound transducers according to In the formula, A ｉｊ is the attenuation of the ultrasound wave as it passes through both sides of the skull, and A ｉ is the skull attenuation of the ultrasound transducer due to the first ultrasound passing through the first side of the skull into the head, and A ｊ is the attenuation of the ultrasound wave by the ultrasound wave exiting the head through a second side of the skull, and k ｉｊ is the inverse cosine of the relative angle of the ultrasonic beam emitted by the first ultrasonic transducer; 5. The ultrasound delivery system of claim 4.

6. 10. The ultrasound delivery system of claim 1, wherein at least one array of ultrasound transducers includes a first ultrasound transducer array and a second ultrasound transducer array, and the head-mounted device includes a frame that couples the first ultrasound transducer array and the second ultrasound transducer array to the head-mounted device on opposite sides of the volume.

7. 10. The ultrasound delivery system of claim 1, wherein the head-mounted device includes a diadem-shaped body having a first opening sized to receive the head and a second opening opposite the first opening that leaves the top of the head exposed when worn.

8. The ultrasound delivery system of claim 1 , wherein the head-mounted device includes an adjustable support mechanism for positioning and supporting the head-mounted device on the head when worn.

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