REFERENCE TO RELATED APPLICATION

[6001] This application claims priority to U,S. Provisional Patent Application Serial No. 62/744,367, filed October 11, 2018, the entire content of which is incorporated herein by reference,

GOVERNMENT SUPPORT

[0002] This invention was made with government support under Grant No. EY029424 awarded by die National Institutes of Health. The Government has certain rights in the invention.

FIELD OF THE INVENTION

[0003] The present invention generally relates to field of interacting with biological tissue using minimally invasive devices, and more particularly using high-resolution devices for dual-mode ultrasonic neural stimulation and imaging.

BACKGROUND OF THE INVENTION

[0004] Real-time interfacing with the brain by monitoring and stimulating brain activity at large scale with high spatiotemporal resolution has the potential to enhance our perceptual, motor, and cognitive capabilities, as well as to restore sensory and motor functions lost through injuries or diseases Currently, the therapeutic utility of brain stimulation in managing numerous neurological and psychiatric diseases has been well understood For example, deep brain stimulation (DBS) ameliorates Parkinson’s disease symptoms ^t¾ , and multiple pilot studies have begun to examine the utility of DBS for neurological conditions such as dystonia ^t4] , epilepsy ^[SI , depression and obsessive^ compulsive disorder ^[7J . A brain-machine interface (BMI) can restore motor control or create actions from thoughts by monitoring brain activity in severely disabled patients suffering from diseases such as spinal cord injury, major amputations, and stroke. In BMIs, neuromodulation can substitute for conventional visual feedback, and provide a direct sensory feedback from artificial limbs into the brain By closing the control Loop in BMIs with neuromodulation, dramatic performance improvements in executing sophisticated tasks with prosthetic devices can be achieved ⁱ⁹¹ .

[0605] In neuroscience research, dynamic mapping of brain circuits by monitoring and modulating brain activity can enhance our understanding of brain functions* and also help researchers to better understand different neurological disorders based on the identification of underlying neural circuits These brain activities can be monitored either directly by recording electrical signals Or indirectly by imaging hemodynamic changes induced by the neurovascular coupling in the vessels surrounding activated neurons ^[ltl . For a neural interface to be an effective research tool, monitoring and modulating brain activity of complex three-dimensional (3D) distributed neural circuits within the entire brain with the high spatiotemporal resolution of several micrometers and milliseconds is often required

1121

[0006] Recently, neuromodulation of the peripheral nervous system (FNS) has also received great attention for its potential to replace prevalent pharmacological approaches with implantable electroceuticals ^[13] . This offers a new class of revolutionary treatments known as bioelectromc medicines that require modulation of neural signaling patterns in visceral organs, such as the lungs, heart, liver, pancreas* kidneys, bladder, gastrointestinal tract, and lymphoid and reproductive organs at the level of nave fibers with diameters of -100 pm and below ^tl4 X Some examples include vagus nerve stimulation for the treatment of epilepsy, inflammatory bowel disease, rheumatoid arthritis, and diabetes, as well as bladder and gastric stimulation ί^H ^ΐ ή Furthermore, the PNS stimulation can also provide sensory feedback to users of prosthetic devices

[0007] Neuromodulation can be achieved with different modalities from pharmacological and chemical methods, which lack specificity and have numerous metabolic requirements, to electrical, electromagnetic, optical, and acoustic methods that have higher specificity, Noninvasive tools, such as transcranial magnetic stimulation (Figure 1 A) mid transcranial direct/altemating-current stimulation (Figure 1 B), suffer from poor spatial resolution (cm and above) P°M221 _r while invasive methods such as electrical (e»g. deep brain stimulation, DBS, in Figure 1C) and optical (optogenetics in Figure ID) stimulation can achieve higher specificity by implanting electrodes and optic-fibers/llght- emitting diodes (LEDs) into neural tissues ^[23H251 . Although optogenetics currently offers unrivaled spatial resolution down to the cellular level, it requires genetic alteration that is risky, if not impossible, in humans. In addition, for current or light delivery, both electrical and optical stimulation requires penetration into the parenchyma of the nervous system that causes sear tissue formation and long-term damage (highly invasive). In the PNS, although cuff or spiral electrodes can deliver stimulation from outside of a nerve, they lack specificity and highly invasive penetrating electrodes are often adopted to stimulate specific fibers inside a nerve ^ί2u 1

[0008] Figure 2A compares the specifications of current methods for brain-activity monitoring ^l27 f Although noninvasive methods such as magnetic resonance imaging (MRI), functional MR!, magnetoeneephalography (MEG), positron emission tomography (PET), and electroencephalography (EEG) provide whale-brain spatial coverage, they suffer from poor spatiotemporal resolution (sub-cm/sub-s and above). High spatial resolution of 1 mm can be achieved with fMRI at the cost of low (sub-minute) temporal resolution and not being portable. In order to improve temporal resolution to the milliseconds scale _* minimally invasive methods such as eleetrocoticography (ECoG) and micro-electecetieography (pECQG) record synchronized potentials close to the cortical surface. However, to achieve higher spatiotemporal resolution (micrometers, milliseconds) inside the brain with a portable system, direct action potential recoding with highly invasive penetrating electrodes is the only viable method that unfortunately cause scar tissue formation and long-term damage ^f2¾ · ^[29 f Figures 7A-7D show different brain stimulation techniques. Figure 7A shows invasive deep brain stimulation (DBS) with electrical currents and Figure 7B shows noninvasive transcrania! direct current stimulation (tDCS). Figures 7€ and 7D show transeranial magnetic stimulation (IMS) and transeranial focused ultrasound (tFUS), respectively.

Ultrasound for Neuromodulation and Hemodynamic imaging

[0009] Low-intensity transcranial focused ultrasound (tFUS) as a noninvasive (external source) neuromodulation modality for both activation and suppression of neural activity in the central nervous system (CNS) and PNS has recently gained more attention due to its higher spatial resolution of sub-cm scale relative to its noninvasive counterparts, Noninvasive. use of low-frequency (0.25-0,65 MHz), pulsed, focused ultrasound for exciting neuronal circuits and motor responses in mouse brains, disrupting electrographic seizure activity in intact mouse brains, stimulating peripheral somatosensory circuits in humans, and modulating the activity of the primary somatosensory cortex in humans has already been shown P ^Q HMl. Recently, successful noninvasive ultrasonic stimulation in mice, rats, cats, rabbits, sheep* monkeys, and humans with different acoustic frequencies within the 0.2-43 MHz range have also been shown (Figure 3A) PSH52] f igure 3 A shows setups for noninvasive ultrasonic neuromodulation on anesthetized / constrained animals and on humans through the skull.

[OOIO! Ultrasound has also been advocated for noninvasive brain activity imaging in small animals with thin skulls due to its higher spatiotemporal resolution of 100 pm and 1 ms in depth relative to its noninvasive counterparts. Functional ultrasound imaging is used for monitoring micro vascular hemodynamics in response to brain activation in rodents.

Using a 15-MHz ultrasound probe, powerful Doppler images of the whole rat brain (2 cm . x 2 cm) have been obtained in 320 ms with a resolution of 100 pm (Figure 3B) ^[531 . Figure 3B shows a setup for performing ultrasonic imaging On tethered animals. As discussed above, ultrasound has already been shown to be an effective noninvasive tool for modulating neural activity with a small latency of tens of ms* and for imaging hemodynamic changes in rodents' brains with spatiotemporal resolution of 100 pm and hundreds of ms. Today's technologies for ultrasonic stimulation are noninvasive, as shown in Figure 3 A, and are applied to the brain tissue through the skull. In addition, brain imaging with ultrasound has only been applied on rodents through removed or thinned skulls. There is no implantable ultrasonic technology for high spatiotemporal resolution (micrometers, milliseconds) neuromodulation or imaging on humans, let alone any dual- modal technology that can support both ultrasonic neuromodulation and imaging. Thus, it is desirable to develop a technology to use ultrasound in humans for neuromodulation and imaging with higher spmtiotenporal resolution to overcome the limitations of the existing technologies and to provide a better solution.

SUMMARY OF THE INVENTION

[Will Real-time interfacing with the central and peripheral nervous systems (CN:S and FNS), such as the brain and nerve bundles in the PNS, by monitoring and stimulating neural activity at large scale with high spatiotemporal resolution (milliseconds, micrometers) has great potentials including but not limited to: 1) enhancing our understanding of the extremely complex nervous system; 2) enhancing our pjerceptual, motor, and cognitive capabilities; 3) restoring sensory and motor functions lost through injuries or diseases; 4) managing numerous neurological and psychiatric diseases (e.g., Parkinson's disease); and 5) replacing prevalent pharmacological approaches with implantable electroceuticals (known as bioeiectronic medicines). Conventional implantable methods for high spatiotemporal resolution neuromodulation or imaging are highly Invasive and require penetration of devices into neural tissue _» which causes longterm damage _* This disclosure outlines a new, minimally invasive, dual-modal implantable technology fin both neuromodulation and imaging at high spatiotempoml resolution (milliseconds, micrometers) and at large scale with ultrasound. A wireless implant, equipped with an array of ultrasonic transducers for both the CNS and PNS, is presented that eliminates the need for penetration into the neural tissue by focusing an ultrasonic beam, Notably, the proposed wireless implant for the PNS utilizes ultrasound for neuromoduiation, imaging, power delivery, and telemetry. In addition, microscopic ultrasonic (pUS) stimulation and imaging is outlined, in which sub-millimeter sized ultrasonic transducers are implanted into the neural tissue (as opposed to sitting on the surface of the neural tissue) to deliver stimulation locally with high precision and also to image the neural acti vity of the neural tissues.

[0012] Similar to Figure 2A, Figure 2B compares key specifications of different neuromoduiation approaches in terms of their spatial coverage, spatial resolution and invasiveness. Noninvasive tools, such as transcranial magnetic stimulation (TMS) and transcranial direct and alternating current stimulation (tDCS, tACS), suffer from poor spatial resolution of centimeter (cm) scale and limited depth of penetration [2]. Invasive methods _» such as electrical and optical stimulation, can achieve better spatial resolution through the implantation of electrodes and optic-fibers^ight-emitting-diodes (LEDs) into neural tissue at the cost of tissue inflammation and damage (highly invasive) [29].

[0013] Low-intensity transcranial focused ultrasound (tFLJS) as a noninvasive neuromoduiation modality for both activation and suppression of neural activity has recently gained more attention due to its improved spatial resolution of sub-cm scale [30]. The tFUS is often practiced at low sonicatibn frequencies (f _P , the frequency at which the transducer is driven) of sub-MHz (i.e., poor spatial resolution) due to the large skull attenuation at high frequencies, but successful ultrasound stimulation atj^s up to 43 MHz has also been reported. [0014] The concept of microscopic ultrasound stimulation (pUS) further improves the spatial resolution and coverage of ultrasound stimulation, hi pUS, either an electronically phased array of ultrasound transducers Or several millimeter (mm>sized focused transducers can directly be placed on the brain surface with partially removed shill (or over thinned skull) as show in Figure 4C or sub-mm-sized transducers can be implanted inside the bran tissue to deliver a focused ultrasound pressure to the neural target.

[0015] Unlike tFUS, in which ultrasound travels through the lossy skull medium with different acoustic impedance compared with the soft tissue demanding low frequency operation (poor spatial resolution), in the pUS acoustic energy is steered mid delivered directly to the neural target within a more homogenous medium. Therefore, neuromodulation with finer spatial resolution at a larger seale can be achieved with higher energy efficiency, defined as the transducer generated acoustic intensity at the neural target divided by the transducer input electrical power. In pUS, high energy efficiency is key in reducing the required electrical power that could consequently enable the pUS system to be wearable or implantable. Despite all: these advantages of the proposed pUS, it is more invasive than tFUS as it requires device implantation. However, pUS is still less invasive compared to its electrical and optical counterparts, because ultrasound transducers in pUS can be placed tin the brain surface outside the parenchyma.

[0016] A key element in both tFUS and pUS is the ultrasound transducers) that converts input electrical power into acoustic pressure, The transducer design can highly affect the stimulation specifications such as spatial resolution and energy efficiency. Most of the prior works have mainly focused on broadband (with low-quality factor, Q) ultrasound transducers with pulse excitation which are frequently used in ultrasound imaging applications. However, the proposed pUS generally requires a sinusoidal excitation (Figure 4D) as well as high energy efficiency (to be portable) that should be achieved with narrowband ultrasound transducers with high Q ,

[0817] Until now, most tFUS experiments have only been focused on studying the impact of ultrasound on neural tissue, investigating optimal tFUS soni cation parameters and studying interactions between acoustic beams mid brain tissues using commercially available bulky ultrasound transducers. Unfortunately there has hem no comprehensive study (in a quantitative manner) on tile impact of transducer dimension, focusing, acoustic matching, backing materials and frequency on the generated acoustic beam profile that relates to the stimulation spatial resolution and energy efficiency.

[6018] In this disclosure, the pUS transducer array in Figure 4C is simplified to a single mm-sized transducer (either focused or With natural focus), made from lead zirconate titanate-5 A (PZT*5A); This disclosure discloses a comprehensive quantitative evaluation of the performance of mm-sized ultrasound transducers (in both water and sheep brain phantom medium) in terms of spatial resolution and energy efficiency in the context of ultrasound neuromodulation that could provide designers of such systems with insight in developing novel electronics and in Vivo experiments. This disclosure introduces a new figure-of-merit (FoM) for pUS as the maximum acoustic intensity to input-power (at electrical port) ratio if PR) related to the energy efficiency.

[6619] A method of neural stimulation and imaging of nervous system of a subject according to this disclosure includes die steps of providing an interface device operable to generate an ultrasonic beam for neuromodulation and imaging of a targeted neural structure of a subject; implanting the interface device in the subject; and providing and disposing an external coil array over toe targeted neural structure of toe subject* wherein the external coil array is wirelessly powering and opmmimicating with toe interface device. Another method further includes the steps of providing toe interface device comprising at least rate ultrasonic transducer operable to focus an ultrasonic beam on toe targeted neural structure of the brain of toe subject for neuromodulation and imaging of the targeted neural structure, and/or comprising at least one sub-millimeter sized ultrasonic transducer operable to generate an ultrasonic point-source near a targeted neural structure of toe brain of the subject for microscopic ultrasound neuromodulation and Imaging of toe targeted neural structure; implanting toe at least one ultrasonic transducer in a subdural region located over a brain surface and/or the at least one sub-millimeter sized ultrasonic transducer inride a neural tissue of the brain of toe subject; and providing and disposing an external coil array over a skull of toe subject, wherein toe external coil array is wirelessly powering and communicating with toe interface device using an inductive link.

[6620] An alternate method further includes the steps of proriding toe interface device comprising at least one ultrasonic transducer operable to focus an ultrasonic beam on a targeted neural structure of a peripheral nervous system (PNS) of toe subject for neuromodulation and imaging of the targeted neural structure, and/or comprising at least one sub-millimeter sized ultrasonic transducer operable to generate an ultrasonic point- source near a targeted neural structure of peripheral nervous system (FNS) of the subject for microscopic ultrasound neuromodid ation and imaging of the targeted neural structure; implanting the at least one ultrasonic transducer over a nerve bundle of the FNS without any penetration into a parenchyma of the FNS, and/or the at least one sub-millimeter sized ultrasonic transducer in a nerve bundle of the FNS of the subject; and providing and disposing an external coil or ultrasonic transducer array over the skin of the subject that is covering the implanted interface device, wherein the external coil or ultrasonic transducer array is wirelessly powering and communicating with the interface device using an inductive or ultrasonic link.

[0021] An implantable neural stimulation and imaging system for a central nervous system (CMS) of a subject according to this disclosure may comprise an interface device configured to be implanted in a subdural region located over a brain surface, the interface device comprising at least one ultrasonic transducer operable to focus an ultrasonic beam on a targeted neural structure of the brain for neuromodulation and imaging of the targeted neural structure; and an external coil array disposed over a skull ofthe subject, the external coil array configured to wirelessly power and communicate with the interface device using an inductive link,

[0022] Another implantable neural stimulation and imaging system for a peripheral nervous system (FNS) of a subject according to this disclosure may comprise an interface device configured to be implanted over a nerve bundle ofthe FNS without any penetration into a parenchyma of the FNS, the interface device comprising at least one ultrasonic transducer operable to focus an ultrasonic beam on a targeted neural structure of the FNS for neutomodulatiem and imaging of the targeted neural structure; and an external coil or ultrasonic transducer array disposed over the skin of the subject that is covering the implanted interface device, the external coil or ultrasonic transducer array configured to wirelessly power and communicate with the interface device using an inductive or ultrasonic link.

[0O23[ An implantable neural stimulation and imaging system for central nervous system (CNS) of a subject according to this disclosure may comprise an interface device configured to be implanted inside a neural tissue of a brain of the subject, the interface device comprising at least one sub-millimeter sized ultrasonic transducer operable to generate an ultrasonic point-source near a targeted neural structure of the brain for microscopic ultrasound neuromodulation and imaging of the targeted neural structure; and an external coil array disposed over a skull of the subject, the external coil array configured to wirelessly power and communicate with the interface device using an inductive link.

[0024] Another implantable neural stimulation and imaging system for peripheral nervous system (PNS) of a subject according to this disclosure comprises an interface device configured to be implanted in anerve bundle of die PNS of the subject the interface device comprising at least one sub-millimeter sized ultrasonic transducer operable to generate an ultrasonic point-source near a targeted neural structure of the PNS tor microscopic ultrasound neuromodulation and imaging of the targeted neural structure; and an external coil or ultrasonic transducer array disposed over the skin of the subject close to the interface device^ the external coil or ultrasonic transducer array configured to wirelessly power and communicate with the interface device using an inductive or ultrasonic link.

[0025] In some embodiments, the interface device is driven at a high frequency (1-40

MHz range) with a continuous, pulsed or sinusoidal carrier wavefbtm, the carrier waveform being amplitude-modulated with a lower frequency (e,g. in a KHz range) or being reconstructed with ¾ train of sharp pulses with varying amplitudes. The pulsed carrier waveform may have a variable number of cycles, pulse repetition frequency and/or duration. The non-invasive embodiment may have the at least one ultrasonic transducer having a maximum thickness of 0-5 mm, These embodiments may have tile interface device comprising at least two ultrasonic transducers that are disposed in either stacked or side-by-side manner. The ultrasonic link in some embodiments is based on a ultrasonic harmonic modulation (UHM) technique.

[0026] In some embodiments, the at least one ultrasonic transducer comprises piezoelectric materials (e.g. lead zirconate titanate, PZT). The at least one transducer in other embodiments is operable to focus the ultrasonic beam having a frequency in a range of 1 MHz to 40 MHz, While in other embodiments, the at least one transducer has a focal length in a range of 1 mm to 50 mm, in alternate embodiments, the at least one sub- millimeter sized ultrasonic transducer is operable to generate the ultrasonic point-source having a frequency in a range of 0.5 MHz to 10 MHz. In yet other embodiments, the at least one sub-millimeter sized ultrasonic transducer has a focal length in a range of 0.1 mm to 1 mm The interface device may comprise an array of the at least one ultrasonic transducers to focus on the targeted neural structure. In some embodiments, _: the interface device may comprise an array of the at least one sub-millimeter sized ultrasonic transducers to focus on the targeted neural structure,

[0027] Some embodiments of die implantable neural stimulation and imaging system may further comprise another interface device configured to be implanted inside the neural tissue of tiie brain of the subject, the interface device comprising at least one sub- millimeter sized ultrasonic transducer operable to generate an ultrasonic point-source near a targeted neural structure of the brain for microscopic ultrasound neuromodulation and imaging of the targeted neural structure; and the external coil array configured to wirelessly power and communicate with the another interface device using the inductive link.

[0028] Yet other embodiments of the implantable neural stimulation and imaging system may further comprise another interface device configured to be implanted in the nerve bundle of the PNS of the subject, the interface device comprising at least one submillimeter sized ultrasonic transducer operable to generate an ultrasonic point-source near a targeted neural structure of the PNS for microscopic ultrasound neuromodutation and imaging of the targeted neural structure; and the external coil or ultrasonic transducer array configured to wirelessly power and communicate with the another interface device using an inductive or ultrasonic link.

BRIEF DESCRIPTION OF THE DRAWINGS

[0029] Figures 1A-1D are images showing examples of current neuromodulation technologies;

[0030] Figures 2A-2B are charts comparing current methods for brain activity monitoring;

[0031] Figure 3A is an image of a current setup for noninVasive ultrasonic neuromodulation on anesthetized / constrained animals and on a human through the skull;

[0032] Figure 3B is an image of a current setup for ultrasonic imaging on tethered animals;

[0033] Figures 4A-4B are images of a proposed implantable dual-modal technology; [0034] Figure AC is a conceptual schematic ©f an embodiment of a microscopic ultrasound stimulation system;

[0035] Figure 4D is a graphical representation of a somfication pattern of file embodiment shown in Figure 4C;

[0036] Figures 5 A-5B are perspective views of an embodiment of a transducer,

[0037] Figure 5C is a transmission line model of an embodiment of a piezoelectric transducer;

[0038] Figures 5D are images of fabricated transducers according to an embodiment of this disclosure;

[0030] Figure 5E is an image of a measurement setup used to measure the acoustic pressure generated by an ultrasound transducer;

[0040] Figure 5F is aschematic of a simulation setup in CQMSOL used to findelectrical and acoustic characteristics of the ultrasound transducers;

[0041] Figures 5G-5H are graphs showing measured transient voltage;

[0042] Figure 6 is a simplified block diagram of a proposed dual-model wireless ASIC for neuromodulation and hemodynamic imaging;

[0043] Figures 7A-7D are images showing different brain stimulation techniques;

[0044] Figure 8 is conceptual diagram of proposed trasnducers implanted over / inside tile brain tissue;

[0045] Figures 9A-9B are simplified acoustic beam profiles of an disc-shaped ultrasound transducer;

[0046] Figures 10A-1OC are images of simulated acoustic intensify profiles of a transducer;

[0047] Figures 1 IA-11B are graphs showing simulated normalized intensify of a transducer,

[0048] Figures 12A-12B are graphs showing simulated normalized intensify of a transducer,

[0049] Figures 13A-13B are graphs showing simulated normalized intensify of a transducer;

[0050] Figures 14A-14D are graphs and images of a beam profile showing a comparison between measured and Simulated characteristics of a transducer;

[0051] Figures 15A-15C are graphs and images of a measured beam profile; [0052] Figures 16A _* 16C are graphs and images showing an impact of a hacking layer on the acoustic beam profile;

[0053] Figures 17A-17D are graphs and images showing an impact of t on the acoustic beam profile;

[0054] Figures 18 A- 18B are graphs and images showing, an impact of Do on the acoustic beam profile;

[0055] Figure 19 shows a measured acoustic beam profile for another transducer;

[0056] Figures 20A-2QE are graphs and images showing an impact of beam focusing and acoustic matching;

[0057] Figure 21 is schematic of an experiment setup used to measure an impact of a sheep brain phantom;

[0058] Figures 22A-22D are images comparing measured normalized acoustic beam profile;

[0059] Figures 23A-23S are schematic representations of an embodiment used for testing on rodents; and

[0060] Figure 24 is a flow chart of a method according to this disclosure.

DETAILED DESCRIPTION OF THE INVENTION

[0061] The proposed implantable technology, which can be realized in two ways as shown in Figures 4A and 4B, may utilize one or an array of ultrasonic transducers in one or multipie implants distributed over the brain/nerve surface or implanted into the brain/nerve tissue for neuromodulation, imaging» or both simultaneously. These minimally invasive implantable ultrasonic transducers in Figure 4A, for neuromodulation» focus an ultrasonic beam towards the targeted neural tissue(s) using beamformihg techniques. In Figure 4B, the sub-millimeter sized ultrasonic transducers implanted at the vicinity of the neural targets) act as point sources and generate ultrasound to stimulate the nearby neural tissue locally. These ultrasonic transducers in Figures 4A and 4B can be driven with a continuous or pulsed waveform or a sinusoidal carrier at high frequency (several MHz) either amplitude-modulated with a lower frequency of hundreds of kHz, or reconstructed with a train Of sharp pulses with varying amplitudes. In pulse-based sonication, the number of cycles, pulse repetition frequency, and duration may be variable. In imaging the brain activity, the number of transducers as receiver or transmitter, and number of pulses in each firing, may be variable.

10062¾ As noted above, certain embodiments of the present invention provide an implantable technology for both the central nervous system (CNS) and peripheral nervous system (PNS), as shown in Figures 4A and 4B, that could utilize one or an array of wirelessly powered ultrasonic transducers for dual-modal neuromodulation and neuroimaging with high spatiotemporal resolution (milliseconds, micrometers) at large scale. Proposed implantable technology can be realized in two ways. In Figure 4A (first method), this technology 100 is minimally invasive, because implants 120 are located over the brain surface 170 (subdural) or a PNS nerve bundle 145 of a subject 10 without penetration into the parenchyma of the nervous system. Wirelessly powered implants, using an inductive link for the brain implants 120 and ultrasonic links for the PNS implants 125, include one or an army of ultrasonic transducers to focus an ultrasonic beam 160 on the targeted neural structure 150. An array of these implants 120/125 may be distributed ova ^* the bain 170 or nerves 145 for large-scale neuromodulation and imaging.

[0063] Proposed implantable dual-modal technology for both neuromodulation and imaging in both the CNS and PNS as shown in Figure 4B (second method) is called hereafter microscopic ultrasound (pUS) stimulation and imaging 200. In this method 200, sub-millimeter sized ultrasonic transducers 220 are implanted into the neural tissue 170 (brain or nerve parenchyma) to deliver stimulation locally with high precision and also image the neural activity. Since each sub-millimeter sized ultrasonic transducer 220 acts as a point source, it can stimulate and image neurons in the vicinity of the transducer with high precision. Bach implant 220 can include one or an array of ultrasonic transducers. An array of these implants 220/225 may be distributed over the brain 170 or PNS nerves 145/155 for large-scale neuromodulation and imaging.

[0064] An external coil array for wireless power / data 110/210 is disposed covering the implants 120/220 over the skin 130; of the skull 140 of the subject 10. Similarly, an external coil array for wireless power/date 115/215 is disposed covering the implants 125/225 over the skin 135 above the nerve bundle /nerve fiber 145/155 of the subject 10.

[0065] Figures 5A and 5B show simplified diagrams forembodiments of the inductively powered brain implants and ultrasonically powered/communicated PNS implants, respectively, with an array of ultrasonic transducers (for method-1 in Figure 4A). Figure 5A shows a proposed miniaturized wireless implant (method- 1) 500 for brain stimulation and imaging that utilizes inductive power delivery and ultra-wideband (UWB) data communication. The implant 500 is embedded in a coating 510 and comprises capacitors 520 _» UWB antenna 530 and wire coil 550. Figure 5B shows a proposed miniaturized wireless implant 120 for stimulating and imaging fibers inside the PNS nerves that also utilize ultrasound lor power delivery and data communication, wherein the front size of this implant is the same as that of Figure 5A with an array of ultrasonic transducers 120. These implants can also have only one transducer for a fixed focus point. The FNS implants use ultrasound for neuromodulation, imaging, power delivery, and data telemetry for the first time. A piezoelectric material is integrated as a 2D array 120 onto an application-specific integrated circuit (ASIC) 540. For the brain implant, a wire-wound coil 530 for power recovery and two off-chip capacitors 520 are added for providing high» voltage regulated supply and resonance (Figure 5A). The implant may have a maximum thickness of 0,5 mm after biocompatible coating 510. For the PNS implant, two stacked ultrasonic transducers 570/580 for power recovery and data transmission may be integrated on the backside of the ASIC 540 using fluough-silicon-via (TSV) technology (Figure 5B). The front side of the ASIC 540 could be similar to that of Figure 5A. The trasnducers 570/580 are wirebonded 560 with the ASIC 540.

[0066] In order to achieve both neutomodulation and imaging within the 3D structure of neural tissue, ultrasonic beamforming with a 2D transducer array is used in this technology

(Figure 4A), In stimulation, the array could only transmit a pre-defined pattern of focused ultraround to the targeted structure. There are several approaches for transmit and receive beamforming, among which the delay-and-sum is the most simple and computationally inexpensive method at the cost of relatively large side lobes ^[541 . To attenuate the ride lobes, transducer elements may be weighted based on their position (called apodization) by exciting them with different voltage amplitudes ^t55} . In imaging hemodynamics related to brain activities, the transducer array may operate in both transmit and receive modes, and the choice of beamforming method can impact image contrast, spatiotemporal resolution, spatial coverage, and system complexity (power consumption, data rate, area). Beamforming may be applied in transmit, receive, or both modes by exciting and then sensing from different numbers of transducers using techniques such as Doppler imaging, synthetic aperture, dynamic focusing* linearly constrained minimum variance, and spatial and frequency compounding, with tradeoffs between resolution, contrast, frame rate, spatial coverage, and system complexity ^I35 i Any beams reflected by the tissues may be received by the transducers. Any beamfonnmg pattern that is fast and provides high- contrast enough to detect blood volume changes in typically 1 s in small blood vessels may be used in this implant

[0667] For pUS stimulation and imaging (Figure 4B), the sub-millimeter sized ultrasonic transducer is implanted (or an array of them are implanted) within the neural tissue and therefore, they locally deliver ultrasound to the target for both stimulation mid imaging. The ASIC 540 in Figures 5A and 5B can also be used to control each transducer in pUS stimulation and imaging. Since these transducers could be miniaturized to submillimeter scale, there is minimal damage to the neural tissue upon implantation. In addition, they act as point sources and locally stimulate the neurons at their vicinity with high precision.

[0668] In this disclosure, disc-shaped piezoelectric transducers for pUS are studied* but a similar theory can be generalized to other types of transducers. For a disc-shaped transducer, there are two strongly excited vibration modes, named thickness extensional (TE) and radial or planar expander (PE) modes, as well as several weakly coupled modes near TE and PE modes. For large aspect ratios (Dofi% where Do and fare the outer diameter and thickness of the transducer (piezoelectric material), respectively, the TE mode is dominant, and the transducer only shows a piston-type displacement Figure 5C shows the Krimholtz, Leedom, and Matthaei (KLM) model of a piezoelectric transducer, in which the transducer is represented by a transmission line tapped at its center. The transducer is driven through the input electrical port. An ideal transformer with the turn ratio of F: 1 models the electrical to mechanical coupling where F can be found from, [0669]

[0676] KT and Z _c are the electromechanical coupling factor and acoustic impedance, respectively. represents the transducer clamped capacitance where a* and A are the damped dielectric constant and the transducer cross-section area. In (1), wo = icvjf is called the half-wavelength resonant frequency in the TE mode where v _a is the stiffened acoustic velodty. The KLM model also includes a capacitor in series with the transformer, (negligible effect, related to the transducer dimension and Kr. Within the transformer secondary side, representing the mechanical part, two generated acoustic waves pass through two transmission lines with //2 length, Z _c characteristic impedance and v _a sound velocity to reach the front (propagating waves in tissue) and back acoustic ports, interfacing with transducer front and hacking materials with Zp and ZB acoustic impedances, respectively _* When ZF and ZB are smaller than Z _c , the transmission line has a resonance close to <¾>». The input electrical impedance at e>e can be found from,

[0071]

[0072] For the fabrication of transducers discussed below, PZT-5A was chosen as the piezoelectric material due to its high electromechanical coupling, which is key in maximizing acoustic pressure and I ² PR. In order to quantitatively study the impact of a wide variety of transducer design parameters, such as the dimension (Z¾* /), f _p , backing material (PCB-backed, air-backed), beam focusing and acoustic matching, nine sets of transducers (US1-US9) wee fabricated (specifications summarized in Table I below) at frequencies of 22-9.56 MHz,

[0073] Figure 5D shows some examples of the fabricated transducers mounted within their assemblies. Figure 5D(a)(e) shows l/<¾ and f/S^ transducers with PCB as the backing material and piezoelectric (PZT) material dimension of D ₀ = 6,8 nun, t = 0,75 mm and A, = 2,8 mm, t ~ 0,3 mm, respectively. The PZT discs were mounted on a 1.5 mm thick PCB using conductive epoxy, and their top silver plate was wirebonded to a PCB pad. The transducers were connected to SMA connectors using a pair of AWG28 wires, Then, their PCBs were placed inside a custom-made 3D-prmted plastic holder. Finally, to provide electrical isolation and protect wirebonds, the transducer surface was coated by Sylgard- 184 (The Dow Chemical Company, Midland, MI) which has similar acoustic properties to water (cannot be used for acoustic matching).

[0074] Figure SD(b) shows the air-backed transducer with the FZT material dimension of Do - 6.8 mm and t = 0.75 mm (£/&). The PZT disc top mid bottom plates were soldered to a pair of AWG28 wires, and then connected to a SMA connector, A 3D-printed plastic holder with a 6.8 mm diameter hole at its center was fabricated to house the transducer and provide air as the backing material. The PZT disc was placed inside the hole and held still by gluing its perimeter to the holder, The other mid of the hole was covered by epoxy to create electrical isolation, thus providing air as the PZT backing material. Similariy, the transducer surface was covered by Sylgard-184. The transducers in Figure 5D(a) to 5D(c) are unfocused but feature a natural focal zoiie as shown in Figure 9B.

[0075] Figures SD(d) and 5D€ shows fabricated focused transducers (USs, US9 ) with a PZT dimension of D _a - 5.8 mm and t - 1 mm and focusing lens materials of EPOTEK- 301 (Epoxy Tech., _; Billerica, MA) and EP0TEK 301+Alumina (AI2O3), respectively

(these materials also provide acoustic matching). The procedure to fabricate these two transducers is as follows. 1) The PZT disc was mounted on a PCB using conductive epoxy. 2) The PCB was placed inside a brass housing to hold the Ions material.3) A steel bearing _: ball with a 10 mm diameter was used as a mold to shape the lens surface (create desired curvature). The bearing ball was held by a 3-axis translation stage (NRT-100, Thorlabs, Newton, NJ) in a way that the distance between the PZT disc surface at its center and the ball surface was - 0.25 mm (2; ultrasound wavelengdi inside the lens, $ = 2.47 MHz) to provide acoustic matching. And finally, 4) The lens material was poured inside the brass housing to fill the space between the PZT disc and the bearing ball. Before pouring the lens material, the ball was sprayed with a mold release agent (Ease Release 200* Mann Formulated Prods., Gillete, NJ), facilitating the ball removal from the transducer when the lens material is cured.

[0076] Figure 5E shows our setup for measuring the acoustic pressure generated by each transducer. The transducers were placed inside a tank filled with deionized (DI) water and were: driven by a function generator with a burst-mode sinusoidal waveform at a pulse repetition interval of 1 ms. All our measurements were done in Dl water to ensure that transducer wires are electrically isolated. Since our goal was to study the acoustic intensity profile of each transducer, they were driven by relatively low ultrasound pressure (tens of kPa), which resulted in no ultrasound scattering. However, in experiments with much higher ultrasound pressure (hundreds of kPa), degassed water may be used to avoid ultrasound scattering due to bubbles and cavitation. To measure the acoustic pressure, the HGL0085 hydrophone (Onda, Sunnyvale, CA) with 85 pm aperture size and 025-40 MHz bandwidth was employed. The voltage across the hydrophone was amplified by a preamplifier with 20 dB voltage gain and 360 pVrm _s input noise. The preamplifier output was captured using a commercial date acquisition system (Razormax 16, GaGe, Lockport,

IL) with a 1 Gsps sampling rate and 16 bits of resolution. Hie measured pressure data was stored in a PC. [0077] To locate the measured pressure with the hydrophone position, a trigger signal from the function; generator was used. The trigger signal was generated 2 ps before each burst Of the input sinusoid. After each 64 consecutive rising edge of the trigger signal <64 ms), the PC sent a command to the 3-axis translation stage to move the hydrophone position with the minimum step size of 50 pm. The stored pressure data was averaged during 64 ms for filtering the high frequency noise.

[0078] Figure 5F show our COMSOL Multiphysics (COMSOL, Burlington, MA) simulation setup used to find the electrical and acoustic characteristic of the ultrasound transducers (US) 120» The transducer front side was covered by Sylgard and was immersed in water with the attenuation coefficient of 0.002 dB/(MHz.cm). The boundaries were modeled with the perfect matching layer (PML). The predefined PZT-5A model in COMSOL was used. The mechanical loss was modeled with the PZT damping factor of 0.02, The Sylgard density and the speed of sound in Sylgard were considered 1500 kg/m ⁵ and 1000 m/s, respectively. The Sylgard thickness was within 1-2 mm for different transducers. In practice, the encapsulation layer thickness should be ~L/4 to maximize the acoustic intensity. But, the Sylgard acoustic impedance is close to that of water, and therefore, its thickness is not an issue. For simplicity, the 3D-printed holder was not modeled in our simulations. In the electrical circuit interface, an AC voltage source and a source resistor (!¾») of 50 f2 were connected to the transducer to measure the driving electrical power and impedance (¾s), as shown in Figure 5F.

[0079] Figures 5G and 5H show the measured hydrophone transient voltage (after preamplifier) when USj was placed at the axial distance of 12.3 mm and 6,9 mm, respectively. The transducer center is considered as the reference point for both axial and lateral distances. The transducer was excited with a burst-mode sinusoid at 2.7 MHz with 12 number of cycles (/½.), As shown in Figure 5G, two different sets of sinusoidal signals can be seen at the hydrophone output. The first set, which coincides With the transducer exciting voltage, was due to the electromagnetic coupling noise between the transducer and the hydrophone. The second set was the actual generated pressure by the transducer.

[0080] Since it takes several cycles for the transducer to reach steady-state vibrations due to its limited bandwidth, Mr should be chosen large enough for the hydrophone voltage to settle down. As the hydrophone was closer to the transducer (axial distance of 6.9 mm) in Figure 5H, the coupling noise interfered with tiie actual generated pressure. For smaller axial distances, the coupling can completely disrupt the actual received pressure waveform. Therefore, measurements in near-field (small axial distances) is quite challenging. To mitigate this, we chose M· as the smallest number of cycles that the transducer vibration could reach its steady state. For example, M _e was 10 and 7 for USi and ti¾, respectively. The axial distances of 6.9 mm and 12.9 mm were primarily chosen to show the coupling noise interference in measuring acoustic intensity at short distances (< 6.9 mm). For neuroscience research applications, these axial distances are optimal for targeting the brain regions of rodents which are within ~10mm of their skull.

[0081] Based on (3), reducing i % decreases the focal length and improves the lateral resolution. We investigated the acoustic intensity profile of sttb-mnnsized transducers (for implantation in brain tissue) using COMSOL. For instance, for a transducer with D _a - 0.5 mm and / = 0.2 mm the focal point was located at the axial distance of < 50 pm. The characteristics of such transduces carinot be accurately measured with our current setup due to die coupling noise. Due to difficulty in fabrication, handling, and characterization of sub-mm-sized transducers, as a first step, in this study we investigated the performance of mm-sized transducers to establish FoM for pUS. These results could help design Of micron-scale ultrasound transducer arrays for pUS (Figure 4C).

[0082] ^' Figure-of-Merit (FoA^ for the Proposed pUS

[0083] For successful ultrasound stimulation, the required acoustic intensity at theneural target should be higher than a threshold, which depends on the sonication frequency and pattern. In pUS, in which the stimulation system could be portable, it is crucial to meet such a threshold requirement with minimal electrical power for driving the ultrasound transducer^ Therefore, we define the maximum acoustic intensity (W/cm ² ) to the input electrical power (W) ratio, termed as fPR (cm ^-2 ), as the FoM for the pUS. This FoM, which directly relates to the stimulation energy efficiency, is used to compare die performance of different transducers

[0084] It is worth noting that the insertion loss (IL) and transducer electrical·*^ mechanical power efficiency (ht) are also the parameters which are often used in the literature as a measure of a transducer sensitivity particularly in imaging. The IL is defined as the voltage inverse ratio of the magnitude of an ultrasound sinusoidal burst, which is emitted by the transducer, to the magnitude of the received reflected echo from a highly reflective plane in parallel with the transducer. The IL is often reported in dB (201ogio). Although IL can be considered as a measure of rfn it does not explicitly relate to tile acoustic intensity as it lacks information about the transducer focusing characteristic. Therefore, both 1L and ht _> while very important, are not the optimal FoM for the pUS compared to fPR. This said, all these three parameters were measured for each transducer as shown in Table Ϊ.

[0886] Microscopic Ultrasound Stimulation Concept

[0087] In the proposed pUS as shown in Figure 8, large external transducers in conventional transcranial BUS systems are replaced with miniaturized ultrasound transducers either implanted inside the brain tissue 170 or placed on the brain surface 170 under the skull 130. For stimulating the regions close to the surface of the brain (within several millimeters), mm-sized transducers 120 with a natural focus are placed on the brain surface (above the target region). To stimulate deep brain regions, sub-mm sized transducers 220, mounted on a flexible thin substrate, with a local ultrasound pressure are implanted inside the brain tissue. The following text discusses the effects of the geometry (diameter and thickness) and f _P of disc-shaped piezoelectric transducers on the stimulation zone. However, similar methodologies can be applied to any transducer shape and material,

[0088] Figure 9A shows a simplified acoustic beam profile of an unfocused disc-shaped transducer with outer diameter df D _a and thickness of i. This beam profile can be divided into near-field and far-field regions, at the boundary of which a natural focal zone is formed with relatively narrower beam and larger acoustic pressure. The near-field region begins from the transducer surface and ends at the focus point, at which the beam reaches its minimum width (~ D</2). Also, the focal length (IV) is defined as the distance between the transducer surface and the focus point Finally, the beam diverges in the far-field region, at which the beam width increases with die distance.

[0089] For the following discussion, we have chosen lead zirconate titanate (PZT) as the piezoelectric material for ultrasound transducer due to its high electromechanical coupling coefficient, which is key in maximizing pressure intensity. There are two main resonance modes in a disc-shaped piezoelectric transducer, called thickness extensions! (TE) and radial or planar expander (PE) modes, as well as several weakly-coupled vibration modes near TE and PE modes [56]. The vibration mode of a PZT transducer is highly dependent on its aspect ratio, defined as DJi. For a large aspect ratio, TE is dominant and the transducer has only a piston-type displacement with high electrical to mechanical efficiency. However, as the aspect ratio is reduced, thickness vibration becomes weaker and non-uniform due to radial mode vibration and its harmonics, resulting in lower electrical to mechanical efficiency.

[0090] The thickness and radial mode resonance frequencies,/* and/*, respectively, of a PZT transducer with large aspect ratio can be found from,

[0091]

[0092] where Afr and NP are thickness and radial mode frequency constants, respectively. It can be seen that both/* and/* are strongly dependent on ZVand t. The focal length (N) also highly depends on the transducer geometry, and can be found for a large aspect ratio from,

[00931

[0094] where $ and v are the sonication frequency and the sound velocity in the medium, respectively.

[0095] Since both (3) and (4) are accurate for a large aspect ratio [56], and the proposed mII8 operates with miniaturized transducers, finite-element method (FEM) simulation tools such as COMSOL Multiphysics (C0MSOL, Burlington, MA) is used to accurately model transducers. Nonetheless, both (3) and (4) can be used for finding initial values for D , [0096] Effects of Transducer Geometry and Sonication Frequency on Acoustic Beam Profile

[0097] In tiiis Section, _: COMSQL simulation results are provided to study the effects of A , /, and ;£ _> on the acoustic beam profile, generated by miniaturized transducers made of PZT-5 A- The minimum required acoustic intensity, which depends on f _p and is within

W/cm ² range, is the key parameter for successful ultrasoimd stimulation [35], In this paper for fair comparison between different cases, normalized acoustic intensity, which is defined as the ratio of the acoustic intensity to the transducer’s input electrical power, is calculated and reported.

[6098] For the pUS model in COMSQL, each miniaturized transducer (US) was located inside the castor oil with an attenuation coefficient of 0.8 dB/cm/MHz to mimic the soft tissue. Figures 10a, 10b, and 10c show the simulated normalized intensity profile for three different transducers with [A = 0.2 mm, _: t = 0.2 mm, f _p = 0.5 MHz], [A = 0.75 mm, t =* 0.2 mm, f _p = 5 MHz], and [A = 4.5 mm, t = 0.2 mm, f _p = 5 MHz], respectively. The transducers in Figures 10a and 10b with A < 1 mm can be considered as examples for implanted (inside the brain) version of the pUS in Figure 8, while Figures 10b and 10c provide examples for pUS from the brain surface. Figure 10A shows the focal intensity maxima around the transducer (i.e. high resolution) at even low f _p of 0.5 MHz. Figures lOB and 10C show focased high-resolution intensity at the axial distance of ~~0.15 mm and ~2 mm, respectively. Increasing D _o and/or f _p can further push the focus point to a larger depth

[0699] The results in Figures 10A-10C show tiie considerable sensitivity of the beam profile, in terms of acoustic intensity magnitude, focal point, side lobes, and near-field effects, to the transducer geometry and f _p . As predicted by (4), by increasing D ₀ aadfp, the focus point moved to larger axial distances^

[00166] Figures 11 A and 11B show the effect of D _o variation of 0.2-1 mm on the normalized intensity for a transducer with t = 0.2 mm, operating at f _p = 0.5 MHz, in axial (zero lateral distance) and lateral (axial distance of 0.1 nun) directions, respectively. For all cases, similar to Figure 10A the focal zone was approximately attached to the transducer surface. Small A (<l mm) results in a short N in (4).

[00161] To consider the spatial resolution, the intensity full width of the half maximum (FWHM) is often calculated [35], Therefore, according to Figures 11A-1 I B minimum FWHM of 60 mih and 150 pm in axial and lateral directions, respectively, can be achieved with Do = 0.2 mm at 100 pm axial distance. This results in a stimulation volume of For comparison, the 0.5-MHz transcranial FUS system in [35] has achieved a measured FWHM of 4.9 mm and 18 mm (stimulation volume of 18*4.9x4.9»432 mm ³ ) in lateral and axial directions, respectively. Therefore, our proposed pUS has reduced the stimulation volume by ~5 orders of magnitude at the same fp. In other words, spatial resolution has been improved from sub-em to sub-mm thanks to the proposed gUS method.

[00102] It can be observed in Figures 11 A-l IB that increasing D ₀ enhances the acoustic intensity at larger axial distances at the cost of increasing FWHM in both axial and lateral resolution» For example, for mm intensity FWHM increased to ~0.3 mm and 0.55 mm in axial and lateral directions, respectively, equivalent to the larger stimulation volume of~0.09 mm ³ ,

[00103] Figures 12A and 12B show the simulated normalized intensity of a transducer (f _p - 0.5 MHz) with different D ₀ and t in axial and lateral directions, respectively. Increasing

Do - 1 from 0,2 mm to 0.5 mm resulted in an increase of the intensity FWHM for both lateral (150 pro to 300 pm) and axial (60 pm to 150 pm) directions. Comparing Figures 1 la-b and 12a-b implies that increasing t for similar ^j ¾ _> can enhance the acoustic intensity with negligible effect on the spatial resolution. For example, the acoustic intensity has increased from 2.6 (Figure 1 IB) to 6.6 (Figure 12B) by increasing t from 0.2 mm to 0.5 mm {Do = 0.5 mm). For the same case, the intensity FWHM in lateral direction has maintained almost constant at 0.32 mm,

[00104] Figures 13a- 13b show the effect Qfj^ {0.5-5 MHz) on the beam profile generated by a transducer with D ₀ = 0.75 mm and / = 0.2 mm. As f _p was increased, according to (2) and Figure 13A the focus point moved to larger axial distances. The intensity FWHM reduced in both lateral and axial directions as f _p was increased (better spatial resolution), At f _P - 5 MHz, minimum intensity FWHM of 75 pm in lateral direction was achieved (Figure I3B). Also as f _p was increased, a significant enhancement in the normalized intensity can be observed in Figures 13A-13B, because^ was closer to fi _r or fir, which are high for small transducers according to (1), Le. more efficient transducer.

[00105] Furthermore, a novel modulation technique^ called ultrasonic harmonic modulation (UHM), for wideband pulse-based ultrasonic communication may be used with the above-discussed embodiments. According to UHM, two narrow pulses with particular amplitudes mid time ^; delays are sent in the transmitter such that their associated ringing ¼ the receiver creates a short dunging by surpassing the ringing of the first pulse with that of the second pulse.

[00106] Smulation ys, Measurement

[00107] To validate the accuracy of our measurements, the simulated and measured characteristics of the transducers woe compared (only U$i simulated results are reported). Table I shows tile specifications and measured electrical and acoustic characteristics of £/&-£/& transducers.

[00108] For each transducer, we report the following measured parameters. 1) electrical impedance measured using a network analyzer (E5071C, Keysight Tech., Santa Rosa, €A), 2) acoustic intensity profile in axial and lateral directions normalized to the maximum intensity in that specific measurement, 3) axial and lateral resolution, defined as halfpower (-3 dB) beam width (normalized intensity reduced to 0.5) in axial and lateral directions, respectively. For axial resolution, the beam width in the ffcr-field region starting at JV is reported, because near-field measured intensities are not reliable. For the lateral resolution, the beam width at N is reported. 4) fPR as discussed earlier.5) /£, which was measured using the similar method in. ft should be noted that the measured It results represent two-way insertion loss and include tile losses introduced by the medium (water), the reflector (air), the divergence of the ultrasound beam, which is the main reason for relatively low IL values in Table I, and measurement setup nonidealities. To exclude such losses, 6) the ratio of the acoustic power to electrical power (ht) was also measured. The acoustic power in ht was calculated by integrating the measured acoustic intensity over a plane in parallel with the transducer surface at the food length (/V).

[00109] Figures 14A to 14D compare measured and simulated results of electrical and acoustic characteristics of C/<¾, Figure 14A shows the USi electrical impedance with the simulated/measured series and parallel resonance frequencies of 2,66/2.6 MHz and 2.85/2.8 MHz, respectively. Figures 14B and 14C show the simulated and measured acoustic beam profiles, respectively, at the parallel resonance frequencies for the same sinusoidal input with 1 V amplitude across USj, Figure 14D compares simulated and measured axial and lateral resolutions extracted from Figures 14B and 14C. The simulated/measured Nmd axial and lateral resolutions were 12/14,25 mm» 15/14,75 mm and 1.3/L9 mm, respectively. For the same input, the maximum simulated and measured acoustic intensity at the ideal tone was 19.6 raW/cm ² and 11 mW/cm ² , respectively. This discrepancy could be due to the differences in the boundary condition (related to reflections), PZT-5A parameters, lateral resolution, and electrical impedance (related to input power) between simulation and measurement setups. Therefore, for fair comparison of transducers, the measured normalized acoustic intensity and flPR is also reported in this disclosure,

[00110] In the following subsections, measurement results are used to study the impact of fp, transducer dimension backing layer, focusing* and acoustic matching on the characteristics of the generated acoustic beam profile.

[00111] Sonication Frequency Impact

[00112] Ideally, f should be selected based on the transduce: mechanical resonance frequency, which is for a disc-shaped transducer in the TE mode. But in practice the backing and matching materials can affect m, hi addition, Co and C‘ capacitors in Figure 5C, which are in series with the equivalent mechanical impedance, appear in the electrical impedance (seen from the input electrical port in Figure 5C), thereby deviating the electrical resonance frequency from cop. Typically, the electrical impedance plot shows the series and parallel resonance frequencies as in Figure 14A. For a transducer with a high ht, these two frequencies merge to wo/2p.

[00113] To show the impact of operation at series vs. parallel resonance, the normalized acoustic intensity of USi was measured at three different frequencies of 2.6 MHz (series resonance, Figure 15A), 2.7 MHz (between series and parallel resonance, Figure 15B) and 2.8 MHz (parallel resonance, Figure 14C). At 2,6 MHz, 2.7 MHz and 2,8 MHz, measured N was 20 mm, 14 mm and 14.25 mm, and the maximum fPR was 0A5 cm ^"2 , 3.1 cm" ² and 6.6 cm ² , respectively. This resulted in 14.7 times improvement in fPR by operating at the parallel resonance (2.8 MHz) compared to the series resonance (2.6 MHz). As shown in Figure 15C, axial and lateral resolution changed at different frequencies. In particular, worse resolution and lower f PR were achieved at 2,6 MHz. Therefore, the beam profile is sensitive to f _P , and one should sweep f _p around the transducer resonance frequencies to find the optimal f _p that maximizes I ² PR. This said, in all transducers we found the optimal j _p at or close to the parallel resonance frequency. [00114] It is worth noting that the optimal/ may not be exactly the same as the transducer center frequency (fc) which is usually found from pulse-echo measurements. In this disclosure, the optimal/ is found when fPR is maximized, while operating a t/ may not necessarily result in the maximum I ² PR. Because, for maximizing f PR the acoustic intensity at the focal zone and the input electrical power should be maximized and minimized, respectively, while at / only the pulse-echo frequency spectrum reaches its maximum. For example, the/· of USi is 2.7 MHz, but fPR is maximum at 2.8 MHz. This is mainly due to the higher transducer impedance at 2.8. MHz, which resulted in lower input power.

[00115] Backing Layer Impact

[00110] Acoustic characteristics of the transducer backing layer not only impacts the generated acoustic intensity but can also affect the transducer electrical impedance Based on (2), Ro (transducer impedance at resonance) is inversely proportional to ZB, and therefore, small ZB (high Re) leads to higher Q, enhancing the IJT. In other words, at snail ZB most of the mechanical wave, which reaches the back acoustic port in Figure 5C, is reflected back towards the front acoustic port, generating higher acoustic pressure in the medium.

[00117] To study the backing material impact, we measured and compared the acoustic beam profile of air-backed USr and PCB-backed US ₂ with similar dimension. The ZB values for air and PCB are ~ 4x 10 ⁴ MRayl and 6.6 MRayl, respectively. As shown in Figure Ϊ6A, at the parallel resonance the air-backed USr achieved higher impedance and sharps phase transition, indicating that its Q is higher. The measured acoustic beam profiles of USr and US2 in Figures 14C and 16B ifp,us ₂ ⁼ 2.8 MHz), respectively, as well as their axial and lateral resolutions in Figure 16C are almost similar, indicating that the beam shape is less sensitive to the backing material, although N is slightly higher for US2 (14.25 mm vs. 16.75 mm). The maximum fPRmd hto£ £$//!$¾ were 6.6Z2.4 cm ² (175% increase in fPR by changing the backing layer from PCB to air) and 36.7/16.3%, respectively, implying that air backing with smaller ZB leads to higher energy efficiency as predicted by the theory. Therefore, unlike acoustic imaging, in which backing materials with high ZB and large damping are preferred (low Q, wideband), in pUS (and tFUS) backing materials with low ZB and small damping (high Q, narrowband) are preferable to increase rthe fPR. [00118] Transducer Dimension (D _o , t) Impact

[00119] The transducer dimension, including £>„ and i _> can significantly affect the transducer performance since mo = xvjt and N in (5) (see below) highly depend on / and Da, respectively. To study the impact of t, the measured characteristics of two PCB-backed transducers with similar Z¼ _> = 4.2 mm and different / of 0.73 mm (E/S*) and 0.4 mm (USs) are compared in Figures 10A-10C, As shown in the measured electrical impedance plots in Figure 17A, the parallel resonance frequency of US* was 3.05 MHz which was smaller than that of USj (5.66 MHz) due to its larger /, Figures 17B and 17C shows the measured normalized acoustic intensity profile of US * and USs, respectively. Figure 17D compares the axial and lateral resolution of these two transducers. The US* md USs achieved different N of 435 mm and 14 mm mainly due to their f _P difference. Higher f _p resulted in larger N . In addition, both transducers achieved similar lateral resolution of -4.1 mm. As discussed in this disclosure, since the focal zone of US* was close to the transducer surface, its measured acoustic beam in Fi gure 17B is somewhat affected by the coupling noise.

[00120] In order to study the impact of D ₀ , we measured and compared acoustic beam profile of two PCB-backed transducers with similar t- 0,75 but different D ₀ of 5 mm ((/&) and 6,8 mm ((/¾) at f _p = 2.8 MHz, The measured electrical impedance of E/S? and U8 ₃ was already shown in Figure 16A, indicating that they have similar parallel resonance frequency of ~2.8 MHz. However, at parallel resonance USs with smaller Do achieved higher impedance (¾), because Ro is inversely proportional to Cp © c A eC D/ based on (2).

[00121] Figures 16B and ISA shows the measured normalized acoustic intensity profile of US ₂ and USs, respectively. Figure 18B compares their axial and lateral resolution. The measured jV for US2 and USs were 16.75 mm and 11.25 mm, respectively, Since N is proportional to Da based on (5), US ₃ achieved larger M The lateral resolution of USz with smaller Do was better than that of USs (1.75 mm vs. 2.2 mm), because beam width is proportional to D _a . Finally, the maximum fPR of EZ¾ and USs were 2.4 cm ² and 3.4 cm ² , respectively. Higher l ² PR of US ₃ for 41.6% could be due to its smaller focal zone and higher Q (see Figure 16 A). Therefore, smaller D ₀ helps to improve lateral resolution at shorter N as well as the energy efficiency to some extent for relatively large D _a ft ratio.

[00122] To study the effect of aggressive size scaling (both D ₀ and /), the acoustic intensity profile of a PCB-backed transducer (E/<¾) with small D ₀ ” 2.8 mm and t = 0.3 mm was measured at its parallel resonance frequency of 9.56 MHz, as shown in Figure 19. The axial and lateral resolution was 6.25 mm and 9.5 mm, respectively, which were significantly improved compared to US _? and US'*. This is mainly due to both decrease in Do and increase \nf _p . The maximum fiPR = 16 cm ² of IBs was also much higher than US _? and IBs, mainly due to smaller D ₀ and resulting improved lateral resolution. This said, pUS at higher f _P often requires larger threshold intensity to stimulate the neural tissue.

[06123] Beam Focusing Impact

[00124] Beam focusing is known to improve the spatial resolution of the generated acoustic beam by a transducer. This could consequently enhance the acoustic intensity at the focal zone, improving the energy efficiency. As described in this disclosure;, for the focused transducers we fabricated spherically shaped acoustic lenses (at the front side) with two different materials, EFO-TEK301 <U%) and EPCXTEIOQl+AUimina (IB?), on top of the PZT disc as shown in Figures 5D(d) and 5D(e), respectively. These FCB-backed transducers had similar dimension (D _a = 5.8 mm, t = 1 mm). For fair comparison, another unfocused (flat) transducer (IB?) with EPO-TEK301 matching layer and similar dimension was also fabricated.' It should be noted that Sylgard-184 was not used in US7.9, because sound speed in the lens material should be higher than that in the medium (water in our experiment) for proper acoustic matching.

[00125] Figure 20A shows the measured electrical impedance of US ₇ -9. The parallel resonance frequency of the unfocused (US?) and focused (US$j>) transducers was 2.2 MHz and 2.47 MHz, respectively, Therefore, the focusing lens slightly changed the resonance frequency. For the focused transducers (USs, »), US$ with EPO-TEK as the lens material achieved higher Q (larger impedance and sharper phase transition at resonance), because by adding Al umina to the EPO-TEK as in US _? , the acoustic impedance of the lens material _· (Z _F ·) increased and led to lower Ro at resonance (Iowa· Q) based on (2).

[00126] Figures 20B, 20C, and 20D shows measured normalized acoustic beam intensity of USr-9 _% respectively. Figure 20E compares the axial and lateral resolution of these three transducers. For US7-9, N was 4.6 mm, 8.5 mm and 7 mm, respectively. While the axial and lateral resolution of unfocused US _? was 12 mm and 1.8 nun, respectively, for focused USs and US? the axial and lateral resolution was very close at 6.5 mm mid ~1 mm, respectively. Therefore, both axial and lateral resolution was significantly improved by ~1.8 times (from 12 mm to 6.5 mm and 1.8 mm to 1 mm) using the focusing lens. Finally, the measured maximum fPR of US7-9 were 2.9 cm ^*2 , 9.3 cm ² and 12.9 cm ^"2 , respectively, indicating that focused i¾$> transducers can achieve much higher energy efficiency (3.2 and 4.4 times, respectively) due to their focusing feature.

[00127] Acoustic Matching impact

[00128] Any mismatch between the acoustic impedance of the piezoelectric material, encapsulation layer (Syigard-184 in USi-e) and the tissue medium can affect the generated acoustic intensity by the transducer due to the acoustic reflections. Ideally, similar acoustic impedance for all these three media is needed. But there is a large difference between the acoustic impedance ofthe and water (Z _m = 1.48 MRayl), mimicking the tissue medium in our measurements. To reduce the acoustic reflections due to the mismatch, the encapsulation layer can be employed as the matching layer as welt For optimal acoustic matching, the thickness and acoustic impedance of this layer should be respectively.

[00129] To study the acoustic matching impact, US9 transducer included both EPQ-TEK and Alumina as both the focusing lens and matching material (Figure SD(e)). Alumina was added to EPO-TEK with the volume fraction of 30% leading to the acoustic impedance of

5.6 MRayl The thickness of this layer at the center of the PZT disc was ~>V4 = 0.25 mm (f _p - 2.47 MHz). For fair comparison, focused US& with similar dimension but without Alumina was used. For USs, only EPO-TEK with acoustic impedance of 3 MRayl, which is lower than optimal provided poor acoustic matching. As shown in Figures 20a - 20e, US _# achieved similar performance to USs with respect to the axial and lateral resolution, but it achieved ~1.4 times higher maximum fPR (12.9 cm ^*2 vs. 9.3 cm ^'2 ) even with lower Q as seen in Figure 2QA. Therefore, acoustic matching is key in maximizing the energy efficiency. With otir current fabrication process, it was hard to accurately control the lens material thickness at 0.25 mm for perfect acoustic matching which could further improve fPR.

[001210] Measurement Results with Bio-Phantom

[00131] In order to study the acoustic beam profile in a more realistic setup, several measurements were conducted with a sheep brain phantom (Carolina Biological Simply, Burlington, NC). hi this section, we only provide the measurement results for USi and operating at low and high j£s of 2.8 MHz and 9.56 MHz, respectively. Figure 21 shows our modified measurement setup with a slice of 4.5 mm thick brain phantom filling the space between the transducer (2 mm away from the PZT disc surface) and hydrophone in the water tank. At the axial distance of 14 mm, the acoustic beam profile at the lateral direction (a plane parallel to the transducer surface) with and without the phantom was measured. The axial distance of 14 mm within the focal zone of both transducers was chosen to avoid any physical contact between our fragile hydrophone and the phantom.

[00132] Figures 22A and 22B show the US] measured normalized acoustic beam profile in lateral direction withoutand with the brain phantom, respecti vely. The lateral resolution of USj, which was 2 mm, was not affected by the phantom, however, the maximum acoustic intensity slightly reduced by ~1.1 dB by introducing the phantom. Figures 22C and 22D show similar results for CZ¾. Since USg was operating at the higher f _p of 9,56 MHz, after introducing the phantom both the lateral resolution and maximum acoustic intensity degraded from 0.85 mm to 0.9 mm and by 5,4 dB, respectively. The attenuation coefficient of the soft tissue is within 0.5-1 dB/(MHz.cm), which matches with our measurements. Therefore, while the shape (spatial resolution) of the acoustic beam profile is less sensitive to the slight inhomogeneity of the phantom medium, the acoustic loss increases at the presence of the phantom particularly at higher frequencies.

[00133] Figures 23 A and 23B summarize the integration and development of the whole pUS system into a portable device for proof-of-concept testing on rodents (e,g. rats). The device could mainly integrate mm-sized ultrasound transducers, which were studied in this paper, and the electronics. Due to the transducer’s design in this work, multiple mm-sized transducers can be placed on the animal’s head to target multiple brain regions. The transducer could be any design in Table I, However, based on the presented study the required electrical power can be reduced using air-backing focusing and acoustic matching techniques.

[00134] Building upon our experience, a miniaturized 3B-printed plastic housing may fabricated to hold the air-hacked PZT material and then add an acoustic lens to achieve focusing and acoustic matching similar to the transducer in Figure 5D(e). A3D-printed plastic pedestal may be surgically fixed onto the animal skull and positioned above the neural target. The pedestal could be filled with ultrasound gel to couple the generated ultrasound by the transducer to the tissue. The skull below tiie transducer can be partially removed or thinned to reduce ultrasound attenuation and reflection.

[00135] The electronics could mainly include a multi-functional chip, which may be designed in a high-voltage CMOS process, that could integrate a power management circuitry to provide enough voltage and power levels required in pUS, a high-voltage driver to drive ultrasound transducers, and a configurable stimulation pattern generator to generate the optimal sotiication pattern as shown in Figure 4D. The chip could be integrated either with the transducers or placed on the animals’ back (and wired to transducers) inside commercially available rodent jackets to distribute the weigh* on the animal body. After successful demonstration of tins technology, we could extend to an ultrasound array and electronic beamforming as shown in Figure 4C.

[00136] For both tFUS and pUS, safety against high ultrasound pressure levels should be considered. The Food and Drug Administration (FDA) has determined limitations on both average and peak of exposed ultrasound intensity to the tissue. The spatial-peak temporal- average intensity (/ _¾ »») should be less than 720 m W/em ² and the mechanical index (Ml), which is indicator of ultrasound peak intensity, should be less than 1.9, The requited acoustic pressure for tFUS is often below FDA threshold levels. Since pUS eliminates the effect of lossy skull, its required ultrasound pressure is less than that of tFUS at similar frequencies,

[60137] A comprehensive study on the acoustic and electrical characteristics of mm-sized piezoelectric transducers for pUS applications is discussed above, The operation and transmission line model of disc-shaped piezoelectric transducers were discussed to establish a basis for study ing their acoustic beam profile. Using the PZT-5A piezoelectric material, nine sets of transducers with different dimensions, frequencies, backing materials, focusing features and matching materials were fabricated. Through comprehensive experimental studies of electrical impedance and hydrophone measurements of these ultrasound transducers, the impact of aforementioned design parameters on spatial (axial and lateral) resolution and acoustic beam intensity (related to energy efficiency) was studied. It was shown that transducer miniaturization, beam focusing, acoustic matching and overall quality factor are critical in improving spatial specificity and energy efficiency of pUS. These transducers could be used m vivo applications. These transducers and theft applications are discussed below in more detail.

[00138] Figure 9B shows the generated acoustic beam profile of a disc-shaped transducer which can be divided into two regions: near field and far field. At the boundary of these two regions, a focal zone appears at which the beam width is relatively narrowest (proportional to Do) resulting in the maximum acoustic intensity. As shown in Figure 9B, even an unfocused disc-shaped transducer has a natural focus. The near-field region begins from the transducer surface and ends at the focal point defining the focal length (N),

100139]}

[00140] where v is the sound velocity in the medium. In near-field region (particularly near the transducer surface), there are several local maxima and minima which are highly dependent on the medium (i,e., difficult to model their exact locations).

[00141] The far-field region begins at the focal point at which the beam starts to diverge, thereby increasing the beam width along the axial direction. As shown in Figure 9B, the acoustic intensity decreases monqtonically in the far-fieid region, and its variation in the lateral direction is less than that in the near field. In far field, the beam profile can be predicted with higher accuracy, which is preferable for reliable Operation.

[00142] The dual-modal (modulation and imaging) ASIC, as shown in Figure 6, in foe proposed implantable technology (Figures 4a-4b) has the unique capabilities of 1) generating a highly focused ultrasonic beam with micrometer resolution for neuromodulation (only for method-1 in Figure 4 A) or driving the implanted sub-millimeter sized transducer (only for pUS in Figure 4B), 2) recovering ultrasonic echoes with high spatidtemporal resolution (micrometers, milliseconds) with a variable spatial coverage (mm ² to cm ² ) for hemodynamic imaging, and 3) receiving wireless power both inductively and ultrasonically and generating a high-voltage supply. Figure 6 is a simplified block diagram of foe proposed dual-modal fully ASIC for neuromodulation and hemodynamic imaging. An on-chip digital beamformef may be implemented to control the delay and amplitude of the signal in each cell of foe array that could consist of an analog delay fine and an efficient tunable-amplitude pulser with high-vohage level shifters for driving the transducer array. The imaging block of foe ASIC may either sequentially focus foe ultrasonic beam in tire transmitter and measure echoes of each firing from one transducer, or fire with one transducer in foe transmitter and receive echoes from multiple transducers in the receiver. A wideband transmitter, such as impulse radio ultra-wideband (1R-U WB), transmits foe imaging data to an external processor for reconstructing images with on-chip UWB loop antennas.

[00143] For wideband data transmission in PNS implants, pulse- or carrier-based modulation technique can be used. As shown in Figure 5B, two stacked or side-by-side ultrasonic transducers can be used in the implant for power recovery and data transmission. For pulse-based commumeation, sharp pulses are transmitted to create a short ringing in the receiving ultrasonic transducer.

[00144] Figure 24 shows a flow diagram of a method 2400 of neural stimulation and imaging of nervous system of a subject according to this disclosure. The method 2400 includes the steps of providing an interface device 2410 that Is operable to generate an ultrasonic beam for neuromodulation and imaging of a targeted neural structure of a subject; implanting the interface device in the subject 2430; and providing and disposing an external coil array 2450 over the targeted neural structure of the subject, wherein the external coil array is wirelessly powering and communicating with the interface device. At step 2470 _» the ultrasonic beam is generated by the interface device and the ultrasonic beam is transmitted towards the targeted neural structure for neuromodulating and imaging of the targeted neural structure. As noted above, several approaches for transmitting and receiving the beam (e.g. beamforming) may be used for stimulating (e.g, using beams as point source) and imaging (e.g. detecting blood volume) of the targeted neural structure.

[00145] Step 2490 includes the interface device being wirelessly powered by the external coil array. The interface device is communicating with the external coil array for performing neuromodulation and imaging of the targeted neural structure. For example,, the external array provides signals / commands for performing neuroraodttiation and imaging as well as receives imaging related data from the interface device for further processing. In some embodiments, the imaging related data is transmitted to a central processing unit / control unit for further processing, The central processing unit / control unit may generate the signals / commands for transmission to the external coil array _* If should be noted that a person skilled in the art would be able to use this disclosure for neuromodulation and imaging of the targeted neural structure and determine the miscellaneous requirements needed that are not detailed herein. The steps discussed above need not be followed in the order as shown,

[00146] In another method, step 2410 further includes the steps of providing the interface device comprising at least one ultrasonic transducer operable to focus an ultrasonic beam en tire targeted neural structure of brain of the subject for neuromodulation and imaging of the targeted neural structure, and/or comprising at least one sub-millimeter sized ultrasonic transducer operable to generate an ultrasonic point-source near a targeted neural structure of brain of the subject for microscopic ultrasound neuromodulation and imaging of tiie targeted neural structure; step 2430 includes implanting the at least one ultrasonic transducer in a subdural region located over a brain surface and/or the at least one sub- millimeter sized ultrasonic transducer inside a neural tissue of the brain of the subject; and step 2450 includes providing and disposing an external coil array over a skull of the subject, wherein the external coil array is wirelessly powering and communicating with the interface device using an inductive link.

[00147] In another alternate method* the steps of providing the interface device 2410 further includes comprising at least one ultrasonic transducer operable to focus an ultrasonic beam on a targeted neural structure of peripheral nervous system (PNS) of tire subject for neuromodulation and imaging of the targeted neural structure, and/or comprising at least one sub-millimeter sized ultrasonic transducer operable to generate an ultrasonic point-source near a targeted neural structure of peripheral nervous system (PNS) of the subject for microscopic ultrasound neuromodulation and imaging of the targeted neural structure; step 2430 includes implanting die at least one ultrasonic transducer over a nerve bundle of the PNS without any penetration into a parenchyma of tiie PNS, and/or the at least one sub-millimeter sized ultrasonic transducer in a nerve bundle of the PNS of the subject; and step 2450 includes providing and disposing an external coil or ultrasonic transducer array over skin of the subject that is covering the implanted interface device, wherein the external coil or ultrasonic transducer array is wirelessly powering and communicating with tile interface device using an inductive or ultrasonic link.

[00148] As will be clear to those of skill in the art, the herein described embodiments of the present invention may be altered in various ways without departing from tire scope or teaching of the present invention. As such, this disclosure should be interpreted broadly.

References

[1] M. Nicolelis,“Actions from thoughts,” Nature , 2001.

[2] T. Wagner, A. Valero-Cabre, and A. Pascual-Leone, “Noninvasive human brain stimulation” Annul Rev. Biomed Eng., vol. 9, pp. 527-565, 2007,

[3] V. Gradinaru, M. Mogri, T. Thompson, J. Henderson, and K. Deisseroth,‘Optical deconstruction of paridnsonian neural cirouftty,” Sweatee* vol 324, pp. 354-359, 2009. [4] J. Yianni, P- Bain, N, Giladi, M, Auca, R. Gregory, C. Joint, D. Nandi, J. Stem, R. Scott, and T. Aziz,“Globus pallidus intemus deep brain stimulation for dystonic conditions: a prospective audit,” Mov, Disard , vol. 18, pp. 436-442, 2003.

[5] M. Hodaie, R. Wennberg, J. Dostroysfcy, and A. Lozano,“Chronic anterior thalamus stimulation for intractable epilepsy,” j Epilepsia, vol. 43, pp. 603-608, 2002.

[6] P. Holtzheimer and H.S. Mayberg,“Deep brain stimulation for psychiatric disorders,”

Rev. Neurosci, vol.34, pp.289-307, 201 1.

[7] L. Gabriels, P. Cosyns, B. Meyerson, S. Andreewitch, S. Sunaert, A. Maes, P. Dupont, J.

Gybets, F. Gielen, and H. Demeulemeester,“Long-term electrical capsular stimulation m patients with obsessive-compulsive disorder,” Neurosurgery, vol. 52, pp; 1263-1274, 2003.

[8] A Barbero mid M. Grosse-Wentrup,“Biased feedback in brain-computer interfaces,” J Neuroeng, RehabiL , vol 7, no. 34, pp. 1-4, 2010.

[9] J. Carmena,“Becoming bionic,” Spectrum, vol 49, no. 3, pp. 24-29, 2012.

[10] K. Deisserotb,“Optogenetics Nature , vol. 8, Jan. 201 L

[113 S. Ogawa, T. Lee, A. Kay, and D. Tank,“Brain magnetic resonance imaging with contrast dependent on blood oxygenation,” Awe. Nad. Acad \ Set, USA, vol. 87, pp. 9868-9872, Dec. 1990.

[12] G. Buzsaki,“Large-scale recording of neuronal ensembles,” Nat. Neurosci , vol, 7, no. 5, pp. 446-451, May 2004.

[13] K„ Famm, B. Lift, K. Tracey, E. Boyden, and M. Slaoui,“Drug discovety; A jump-start for electroceutirals, " Nature, vol 496, pp. 159-161, Apr; 2013.

[14] K. Birmingham, _: V. Gradinaru, P. Anikeeva, W. Grill, V. Pikov, B. McLaughlin, P. Pasricha, D. Weber, K. Ludwig, and K. Famm, *¾ioelectronic medicines: A research roadmap,” Nature Rev., vol. 13, pp. 399-400, Jun. 2014.

[15] B. Bonaz, C. Picq, V. Sinniger, J. Mayol and D. Clarencon,“Vagus nerve stimulation: from epilepsy to the cholinergic andti-inflammatory pathway,” Neurogast, Mot, vol. 25, pp.208- 221, Mar.2013.

[16] J, Lee, D. Kim, S. Yoo, H, Lee, G. Lee, and Y. Nam,“Emerging neural stimulation technologies for bladder dysfunctions” M. Neurol J., vol. 19, pp. 3-11, Mar. 2015.

[ 17] G. O’Grady, J. Egbuji, P. Du, L. Cheng, A. Pullan, and J. Windsor,“High-frequency gastric electrical stimulation for the treatment of gastroparesis: a meta-analysis,” World J. Surgery, voL 33, no. 8, pp, 1693-1701, Aug. 2009.

[18] H. Zbu, H. Sallam, and J. Chen,“Synchronized gastric electrical stimulation enhances gastric motility in dogs,” Neuragastroenterol Motif., vol 293, no. 5, pp. 1875-1881, Nov. 2007.

[19] S, Nag and N Thakor, “Implantable neurotechnologies: electrical stimulation and applications,” Med BioL Wing. CompuL,\o\ _> 54, no. 1, pp. 63- 76, 2016.

[20] A. Barker, “The history and basic principles of magnetic nerve stimulation,” Electroencephcdogr. Clin. Neurophysiol Suppl, vol, 51, pp. 3-21, 1999.

[21] T. Wagner, A. Valero-Cabre, and A. Pascual-Leone, “Noninvasive human brain stimulation," Arum Rev, Blamed Eng., vol 9, pp. 527-565, 2007.

[22] A. Javadi, A. Beyko, V. Walsh, and R. Kanai,“Transcranial direct current stimulation of the motor cortex leases action choice in a perceptual decision task,” J. Cognitive Neurosci, vol. 27, pp. 2174-3185, Oct, 2015.

[23] Deep Brain Stimulation Systems, Available Online: http://www.medtronic.com/

[24] E. Boyden,“A history of optogenetics: tbe development of tools for controlling brain circuits with light,” Biology Report, May 2011.

[25] T. Kim, J. McCall, Y. Jung, X. Huang, E. Siuda, Y. Li, J. Song, Y. Song, JL Pao, R, Kim, C, Lu, S. Lee, S, Song, G. Shin, R. Al-Hasani, S. Kim, M. Tan, Y. Huang, F. Omenetto, J. Rogers, and M. Bruchas,“Injectable, cellular-scale optoelectronics with applications for wireless optogenetics," Science, vol. 340, pp. 211 -216, Apr. 2013. [26] Utah Slanted Electrode Array, Blackroek Microsystems Inc.; Retrieved Jul. 2016, Available Online: https y'/commonfmid.nih.gGv/sites/default/files/Blackrocklnfo.pdf

[27] S, Ha, A. Akinin, J. Paris, C. Kim, H, Wans C. Maier, P. Meroier, and G. Cauwenberghs, “Silicon integrated high-density electrocortical interfaces,” Froc. IEEE, vol. 105 _» pp, 11-33, Jan.2017.

[28] M. Yin, D. Borton, J. Aceros, W. Patterson, and A. Nurmikko,“A 100-channel hermetically sealed implantable device for chronic wireless neurosensing applications,” IEEE Treats, Biomd Or. Syst, yol j. pp. 115-128, Apr. 2013.

[29] G. McConnell, H. Rees, A. Levey, C. Gutekunst, R, Gross, and R. Bellamkonda,“Implanted neural electrodes cause chronic, local inflammation that is correlated with local neorodegeneration,” J Neural Eng., vol. 6, p. 056003, Oct. 2009,

[30] W. Tyler, Y, Tufail, M, Finsterwald, M. Tauehmann, E. Olson, and C. Majestic,“Remote excitation of neuronal circuits using low-intensity, low-frequency ultrasound,” Plos One, vol. 3, Oct 2008.

[31] Y. Tufail, A _> Matyushqv, N. Baldwin, M, Tauehmann, J. Georges, A. Yoshihiro, S. Tery, and W. Tyler,“Transcranial pulsed ultrasound stimulates intact brain circuits,” Neuron, vol. 66, pp. 681-694, lint.2010.

[32] Y. Tufail, A. Yoshihiro, S. Pati, M. Li, and W, Tyfer,“Ultrasonic neuromodulatiott by brain stimulation witii transcranial ultrasound,” Nature Protocols, vol. 6, no. 9; pp. 1453-1470, 201 L

[33] W. Legem, A. Rowlands, A. Opitz, T, Sato, and W. Tyler,“Pulsed ultrasound differentially stimulates somatosensory circuits in humans as indicated by EEG and fMRI,” PLOS One , vol. 7, Dec. 2012.

[34] J. Mueller, W. Legpn, A. Opitz, T. Sato, and W. Tyler,“Transcranial focused ultrasound modulates intrinsic and evoked EEG dynamics,” Brain Stint., vol. 7, pp.900-908, Sep.2014

[35] W. Logon, T. Sato, A. Opitz, J, Mueller, A. Barbour, A. Williams, and W. Tyier, “Transcranial focused ultrasound modulates the activity of primary somatosensory cortex in humans^” Nature Nettrosci., vol. 17, pp. 322-333, Feb. 2014.

[36] B. Min, A. Bystritsky, K. Jung, K. Fischer, Y. Zhang, L. Maeng, S. Park, Y. Chung, F. Jolesz, and S. Yoo, “Focused ultrasound-mediated suppression of chemically-induced acute epileptic EEG activity,” BMC Neurosci.,pp. 12-23, 2011.

[37] S. Yoo, A Bystritsky, J. Lee, Y. Zhang, K. Fischer, B. Min, N. McDannold, A. Pascual- Leone, and F. Jolesz, “Focused ultrasound modulates region-specific brain activity,” Neuroimage, vol, 56, pp. 1267-1275, Feb. 2011.

[38] H. Kim, S. Taghados, K. Fischer, L. Maeng, S. Park, and S, Yoo,“Nonmvasive transcranial stimulation of rat abducens nrve by focused ultrasound,” Ultrasound Med. Biol., vol. 38, pp. 1568-1575, 2011

[39] W, Lee, Y. Chung, Y, Jung, I. Song, and S. Yoo,“Simultaneous acoustic stimulation of human primary and secondary somatosensory cortices using transcranial focused ultrasound,” BMC Neuroscl , vol, 17, 2016.

[40] W. Lee, H. Kim, Y. Jung, I. Song, Y. Chung, and S. Yoo,“Image-guided transcranial focused ultrasound stimulates human primary somatosensory cortex,” Scien. Reports, vol. S, pp. 1-10, Mar. 2015.

[41] W. Lee, H. Kim, Y. Jung, Y. Chung, I. Song, J. Lee, and S. Yoo,‘Transcranial focused ultrasound stimulation of human primary visual cortex," Scien. Reports, vol. 6, pp. 1-12, Sep. 2016.

[42] W. Lee, S. Lee, M. Park, L· Foley, E Purcell, H, Kim, K. Fischer, L, Maeng, and S, Yoo, “Image-guided focused ultrasoutidr-mediated regional brain stimulation in sheep,” Ultrasound Med, Mol, vol. 42, pp. 459-470, 2016. [43] T. Deflieux, U. Younan, N. Wattiez, M. Tanter, R. Pouget, and J. Aubry,“Low-intensity focused ultrasound modulates monkey visuomotor behavior” Current Biology, vol. 23, pp. 2430-2433, Dee.2013.

[44] J„ Kubanek, J. Shi, J. Marsh, D. Chen, C, Deng, and J. Cui, "Ultrasound modulates ion channel currents,” Nature Scien. Rep., vo\. 6, pp. 1-14, Apr, 2016.

[45] E. Juan, R. Gonzalez, G. Albors, M. Ward, and P. Irazoqui,“Vagus nerve modulation using focused pulsed ultrasound: potential applications and preliminary observations in a rat,” int J. Imaging SysL TechnoL, vol. 1, pp. 67-71 , Mar. 2014.

[46] H. Back, K. Pahk, and H. Kim,“A review of low-intensity focused ultrasound for neuromodulation,” Biomed Eng. Lett., Jan. 2017.

[47] A. Bystritsky and A. Kerb, *A review of low-intensity transcranial focused ultrasound for clinical applications,” Curr. Behav. Neurosci, Rep., vol 2, pp. 60-66, 2015.

[48] L, Ai, _: J. Mueller, A. Grant, Y. Eryaman, and W. Legon,“Transcranial focused ultrasound for BOLD fMRI signal modulation in humans^ IEEE 2016.

[49] T. Dickey, R. Tych, M. klkrt, J, Loseser, K. Pederson, and P. Mourad,“Intense focused ultrasound can reliably induce sensations in human test subjects in a manner correlated with the density of their mechanoteceptors” Ultrasound in Med Biol., vol. 38, vo. 1 , pp. 85-90, 2012.

[50] E. Mehic, J, Xu, C. Caler, N. Coulson, C. Moritz, and P. Mourad,“Increased anatomical specificity of neuromodulation via modulated focused ultrasound” PLOS One, vol. 9, Feb, 2014.

[51] R, King, J. Brown* and K, Pauly, “Localization of ultrasound-induced in vivo neurostimulation in the mouse model,” Ultrasound Med Biol, vol. 40, no. 7, pp, 1512- 1522, 2014.

[52] M. Menz, O. Oralkan, P. Yakub, and S. Baccus,“Precise neural stimulation in the retina using focused ultrasound," J. Neurosci., vd.33, pp.4550-4560, Mar.2013.

[53] E. Mace, G. Montaldo, B. Osmattski, L Cohen, M. Fink, and M. Tantef,“Functional ultrasound imaging of the brain: theory and basic principles,” IEEE Tram . Ultras. Ferr. Freq, Coni., vol.60, pp. 492-506, Mar, 2013.

[54] B. Van Veen and K. Buckley,“Beamforming: a versatile approach to spatial filtering, ^M /#M ASSP Mag., Apr. 1988.

:[S5] K. Thomenius,“Evolution of ultrasound beamforrners,” IEEE Ultras. Sym., 1996.

[56] M. Meng, and M. Kiani,“Design and optimization of ultrasonic wireless power transmission links for millimeter-sized biomedical implants,” IEEE Trans. Biomed Ctr . SysL, vol. 11 , no. 1, pp. 98-107, Feb.