Magnetoelectric data and power delivery to miniature biodevices with variable amplitudes and waveforms

Abstract

The present disclosure describes new devices, systems, and methods utilizing a magnetoelectric nerve stimulation device with variable amplitude and waveform. Specific aspects of the present disclosure include a magnetoelectric film, a magnetic field generator, and an electrical circuit coupled to the magnetoelectric film. In certain aspects, the electrical circuit includes components configured to modify an electrical output signal generated by the magnetoelectric film. In certain aspects, the electrical circuit is configured to modify the electrical signal to charge a charge storage element, transmit data to an implantable wireless neurostimulator, and provide a stimulation output to an electrode. TIFF2022527987000009.tif84155

Description

[Technical field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application claims priority to U.S. Provisional Patent Application No. 62 / 830,089, filed April 5, 2019, and U.S. Provisional Patent Application No. 62 / 976,051, filed February 13, 2020. The entire contents of each of these applications are incorporated herein by reference.

[0002] A. Field The present disclosure relates to devices and methods that utilize magneto-electrical nerve stimulation devices with variable amplitudes and waveforms, and in particular to devices and methods configured to charge a charge storage element, transmit data to an implantable wireless nerve stimulation device, and provide a stimulation output. [Background technology]

[0003] B. Related Technology There are many challenges to the successful realization of implanted powered devices. For example, a fundamental problem in bioelectronics is the ability to power small devices inside the body. Wires provide efficient power transmission but are a common point of failure and limit device placement. Wireless powering by electromagnetic waves or ultrasound must also overcome obstacles. For example, wireless powering by electromagnetic waves or ultrasound must overcome absorption by the body and impedance mismatch between air, bone and tissue. Also, conventional methods for wirelessly powering neural implants in deep tissue regions of freely moving animals or humans are usually bulky due to large electromagnetic coils or battery packs with external leads. In addition, the ability to provide magneto-electrical charging, data transmission and stimulation to implanted wireless neurostimulators is not provided in existing systems.

[0004] Thus, these and other issues must be addressed for an effective implementation of an implanted powered device. Summary of the Invention

[0005] overview Briefly, the present disclosure provides a system for transmitting data via a magnetic field to an implanted neurostimulator device. In certain embodiments, the implanted device includes electrical circuitry that extracts timing and voltage references from the magnetic field to perform reliable, calibration-free data transmission. In certain embodiments, the circuitry transitions between different operating modes (power transmission, data transmission, stimulation) by monitoring the envelope of the magnetic field, which can be more reliable and energy-efficient than some conventional methods.

[0006] In addition, the present disclosure provides devices, systems and methods that utilize magnetic fields that are largely free from absorption by the body or impedance differences at interfaces between air, bone and tissue. These advantages have led to magnetically powered stimulators based on induction or magnetocaloric effects. However, fundamental limitations in these power transmission technologies have prevented the application of miniature magnetically powered stimulators in many therapeutic modalities and disease models, as such stimulators do not operate in the clinical "high frequency" range above 20 Hz.

[0007] Exemplary embodiments of the present disclosure utilize magnetoelectric materials that provide small-scale energy harvesting from magnetic fields for neural stimulation or to power electronic devices. Exemplary embodiments rely on an alternating magnetic field emitted by a magnetic field driver that modulates a high-frequency (e.g., 20-500 KHz) magnetic field provided by an electromagnetic coil. This magnetic field is at or near the resonant frequency of a magnetoelectric thin film, which can typically be placed a few centimeters away from the coil. The thin film (a laminate of piezoelectric and magnetoelectric materials) can convert the magnetic field into a high-frequency high-voltage electrical signal. Exemplary embodiments attach circuit elements to the thin film to vary the voltage waveform (e.g., rectify and cap the voltage and current to stable levels for a desired application). In specific exemplary embodiments, the electric field from the thin film can be used to directly stimulate excitable cells in vitro, and the thin film can be coupled to electrodes implanted in vivo for stimulation in freely moving subjects.

[0008] Magnetoelectric devices that are remotely powered using magnetic fields can deliver specific electrical signals in a small device package size. This miniaturization is key to making the devices wearable and targeting hard-to-reach areas of the body, including deep brain and peripheral regions. In addition, wireless stimulators help avoid some of the major failure modes of neurostimulators, such as lead migration and contact failure. A further advantage is that the pattern of stimulation can be selected by an externally controlled magnetic field driver to create specific timed pulses of alternating magnetic fields of specific frequencies. A further use of this magnetoelectric voltage control can be to power neural implants in the central or peripheral nervous system for recording physiological variables in addition to electrical stimulation.

[0009] Certain embodiments include four different waveforms that can be useful for stimulating neural tissue depending on the desired application. The various waveforms are created by using the correct number of membranes and their attached circuits and modulating the frequency of the magnetic field appropriately slowly. The waveforms include: (1) high frequency monophasic (single membrane); (2) envelope monophasic (single membrane); (3) high frequency biphasic (double membrane) and (4) envelope biphasic (double membrane or single membrane).

[0010] Exemplary embodiments of the present disclosure can generate a variety of therapeutic neurostimulation by incorporating appropriate magnetoelectric thin films, circuit elements, and magnetic field generator waveforms. For example, the natural resonant frequency of the thin film may not generate a physiological response, but exemplary embodiments of the present disclosure can create lower frequency biphasic pulses within the therapeutic range. Exemplary embodiments can also pulse the magnetic field, thus generating voltage pulses that can directly modulate cellular activity using high frequency resonance. Aspects of the present disclosure can achieve this with one thin film for monophasic voltage signals, or two thin films for biphasic voltage signals.

[0011] In addition, typical voltages reported from magnetoelectric thin films are generally only a few volts at most, meaning that the voltage drops rapidly with the angle of the film and the distance from the magnet. These low voltages are due to the weakness of the alternating magnetic field used to generate the magnetoelectric voltage. At such high frequencies, these strong magnetic fields are difficult to create without a high voltage (roughly kV range) power supply. An exemplary embodiment of the present disclosure uses a custom designed magnetic field driver to generate a high intensity magnetic field using a small system that requires only a few tens of volts, which can be plugged into a common outlet. This configuration can generate a maximum voltage of greater than 50V. To stabilize this stimulus, an LED or Zener diode can be used. This caps the voltage at a value selected depending on the application, typically 1-6V. Thus, the exemplary embodiment can stably output a given voltage even with variations in distance and angle.

[0012] Although power efficiency is lower than some computing technologies, tissue absorption is much lower than other wireless power schemes, allowing the ability to safely power devices deep within tissue. The efficiency of power transfer is angle dependent, therefore the driver coil and device should be aligned for maximum efficiency.

[0013] Examples disclosed herein include relatively straight magnetoelectric material geometries. Other embodiments may include magnetostrictive devices of various shapes or configurations, as well as using different materials for the magnetostrictive and piezoelectric components. Certain embodiments may also include magnetic elements in the thin film to bias the thin film to enhance sensitivity. Embodiments may also include thin films that naturally rectify the voltage waveform. In certain embodiments, the bias magnetic field may be incorporated into either the thin film or the alternating magnetic field circuit. Specific embodiments may include a miniature form factor system that can be fully embedded in a moving target.

[0014] Miniaturized wireless bioelectronic implants can provide a powerful capability to accelerate fundamental biological research and improve clinical therapy for human diseases. Thanks to their small size and wireless interface, these tiny devices have the potential to provide a more non-invasive and longer-lasting interface to tissues such as the brain or peripheral nerves. One of the fundamental challenges in developing these implantable bioelectronics is wireless power and data transmission (WPDT) to devices inside the body. Existing technologies based on radio frequency electromagnetic fields, ultrasound, inductive coupling and magneto-thermal effect cannot meet all of the desired properties of a safe, reliable, stable and efficient WPDT scheme. To address these challenges, we propose to design magnetoelectric (ME) materials and integrate them with CMOS chips to create the first WPDT technology based on magnetoelectric effect. This proposed magnetoelectric "smart dust" technology (sometimes referred to herein as "Mag-Dust" or "MagMote") enables many scientific and medical applications by simultaneously solving three challenges facing miniature bioelectronic implants: low absorption and reflection by tissue and bone; stable performance regardless of changes in angle with and distance to the transmitter; and wearable power / data transmitter.

[0015] Exemplary embodiments of the present disclosure incorporate magnetoelectric (ME) wireless power and data transmission technology for small, highly integrated bioelectronic implants. Exemplary embodiments of the present disclosure achieve significant advances in three technology areas: (1) a new manufacturing strategy to produce miniaturized, highly efficient ME transducers for wireless power and data transmission to implanted devices; (2) a novel low-power CMOS circuit and system design for Mag-Dust, aiming to improve power efficiency and robustness, reduce calibration effort, and increase yield; and (3) a proof-of-principle 1-millimeter cubed programmable ME-enabled wireless peripheral nerve stimulator. This work also produces an integration strategy to achieve high biocompatibility, long-term stability, and safety, as evaluated in in vitro and accelerated degradation tests. In addition, we demonstrate a battery-powered portable magnetic field transmitter to drive and program Mag-Dust.

[0016] Exemplary embodiments of the present disclosure can have broad implications for miniaturized bioelectronic implants. Specifically, exemplary embodiments can solve the fundamental challenge of safe, reliable and efficient wireless power and data transmission to deeply implanted miniature devices, paving the way for minimally invasive and long-term interface technologies for long-term neural stimulation, recording, drug delivery, physiological monitoring, and more. These technologies are expected to benefit numerous transformative and far-reaching scientific and medical applications in understanding the nervous system (especially neuronal and social behaviors), treating disorders, and alleviating pain.

[0017] Bioelectronic implants offer powerful capabilities to accelerate fundamental neuroscience research and improve clinical treatments for human disorders. Common applications include electrical and optical stimulation and inhibition as well as measurement of neural and physiological signals. Traditional battery-powered implants are often too large to be directly implanted at the target site, and the use of wires to connect the battery pack to the sensor / stimulator is a common failure point for bioelectronic devices. In clinical applications, leads present a route for infection [1] and limit the ability of the stimulator to move with the tissue, leading to foreign body reactions, or loss of contact with the target tissue [2, 3]. In scientific experiments using small animals, wires or heavy batteries can limit natural behavior, especially when studying social interactions between multiple animals [4]. Thus, technologies to miniaturize implants and eliminate cords are critical in building next-generation minimally invasive, high spatial resolution, and long-term bioelectronic interfaces to tissues of the central and peripheral nervous systems. These miniature, wireless bioelectronic implants have the potential to enable and advance more than a few transformative clinical and scientific applications.

[0018] One of the fundamental challenges in the development of miniature wireless implantable bioelectronics is to safely and reliably deliver power and data to small devices inside the body. Efficient wireless power and data transmission via electromagnetics (EM) to implantable mm-sized receivers requires GHz radio frequency EM waves. At these frequencies, EM radiation is significantly absorbed by the body. As a result, there are severe limitations on the amount of power that can be safely delivered [5] and the volume of the implant [6-8]. Ultrasound is relatively immune to tissue absorption, but is significantly reflected (attenuated) by acoustic impedance mismatches between air, bone and tissue, making it difficult to deliver sufficient energy through air or bones such as the skull. Optical power transmission is an alternative approach to eliminate bulky receivers [9], but suffers from limited penetration depth and tissue heating. Magnetic fields, on the other hand, suffer little energy loss due to absorption and reflection by the body [10-12]. These advantages have led to magnetically powered implants based on inductive coupling or the magneto-caloric effect. Inductive coupling is a well-developed technology for power transmission in consumer electronics, but is limited by the diameter of the receiving coil. Reducing the coil diameter below 9 mm significantly reduces the received power and makes it more sensitive to the distance and alignment between the transmitter and receiver [13, 14]. Finally, the recently demonstrated magnetothermal stimulation eliminates the receiver size concern, but does not operate in the clinical "high frequency" range above 20 Hz [15, 16], making it unsuitable for many therapies and disease models. Moreover, the required magnetic field strengths of >20 mT cannot be easily produced by portable generators, which limits the ability to use this technology for long-term neuromodulation therapy in freely moving patients or animals.

[0019] In summary, an ideal wireless power and data transmission technology should possess the following properties: (1) low absorption and reflection by the body to safely deliver sufficient energy to implants; (2) controlled and reliable wireless power and data transmission; and (3) a battery-operated, portable transmitter to enable biomedical applications requiring long-term use.

[0020] To address these fundamental challenges, we propose Mag-Dust, a transformative approach to create the first miniature wireless bioelectronic implants that combines magnetoelectric (ME) materials with CMOS ASICs. By coupling magnetostrictive and piezoelectric effects in nanofabricated ME transducers, low-frequency (100 kHz-10 MHz) alternating (AC) magnetic fields can be efficiently converted into alternating potentials to enable wireless power and data transmission with advantages including the fact that low-frequency magnetic fields are not absorbed by the body and therefore can efficiently power deeply implanted devices. Preliminary results have confirmed the ME effect more than 8 cm beneath bone and tissue under a 1 mT magnetic field

[17] . In addition, because acoustic resonances have wavelengths 105 times shorter than electromagnetic resonances of the same frequency, ME technology miniaturizes implanted devices without significantly reducing the voltage generated by the transducer. Furthermore, ME transducers with high voltage transduction efficiency can potentially be driven by portable and lightweight battery-powered magnetic field generators, paving the way for new paradigms of long-term and non-restrictive neuromodulation and recording systems for longitudinal scientific studies in freely moving rodent models as well as future wearable point-of-care disorder treatment and pain relief devices.

[0021] Based on ME-enabled wireless power and data transmission technology, exemplary embodiments of the present disclosure include a fully integrated proof-of-concept millimeter-sized Mag-Dust neurostimulator. In designing and manufacturing the ME transducer, fabrication techniques can be used to create mm-scale ME transducers for wireless power and data transmission based on polyvinylidene fluoride (PVDF) or lead zirconate titanate (PZT) and magnetostrictive alloy (Metglas). Further embodiments may include improving miniaturized ME transducer efficiency by adding magnetic bias. In the case of Mag-Dust CMOS circuit and system design, CMOS circuits and systems can be used to interface with the ME thin film to enable controlled power delivery, calibration-free robust data transmission and programmable stimulation. Exemplary embodiments include a millimeter-sized neurostimulator with biocompatible integration of the ME transducer with the CMOS chip. In vitro testing can be performed to evaluate the functionality and long-term stability of exemplary embodiments in aqueous solutions that approximate conditions in the body.

[0022] Exemplary embodiments of the present disclosure include magnetoelectric (ME) wireless power and data transmission technologies for miniaturized, highly integrated bioelectronic implants. Exemplary embodiments include low power CMOS circuit and system designs for increased power efficiency and robustness, reduced calibration effort, and improved yield. In addition, exemplary circuit technologies provide low power energy harvesters and voltage / current / timing references. Exemplary embodiments include low power CMOS circuit and system designs for increased power efficiency and robustness, reduced calibration effort, and improved yield. 3 The device includes a device configured as a programmable wireless neurostimulator. The integration strategy is studied for biocompatibility and safety, which is evaluated with accelerated degradation and in vitro tests. Exemplary embodiments also include a battery-powered, wearable 1 mT magnetic field transmitter to program and maintain the biphasic stimulation of the proposed Mag-Dust stimulator, covering a wide range of stimulation patterns appropriate for therapy (frequency from 0 to 200 Hz, amplitude from 30 to 1500 μA, and pulse width from 50 to 1000 μs).

[0023] Exemplary aspects of the present disclosure demonstrate that magnetoelectric materials, applied for the first time to bioelectronic devices, enable miniature magnetically powered neurostimulators operating at clinically relevant high frequencies. As an example, it is shown that ME neurostimulators can effectively treat symptoms of Parkinson's disease models in freely moving rodents. We also show that ME-powered devices can be miniaturized to a size smaller than a grain of rice while maintaining effective stimulation voltages. These results suggest that ME materials are excellent candidates for wireless power delivery to enable miniature neurostimulators in both clinical and research applications.

[0024] Wireless neurostimulators have the potential to provide a less invasive and more durable interface to brain regions and peripheral nerves compared to battery-powered or wired stimulators. Indeed, wires are a common failure point for bioelectronic devices. Percutaneous wires present a route for infection [1], and implanted wires also limit the ability of the stimulator to move with the tissue, leading to foreign body reactions or loss of contact with the target tissue [2, 3]. In addition, especially for peripheral devices, long-term stress and strain on the wire can lead to failure of the wire itself or its connection to the stimulator [4]. In small animals such as rats and mice, the wires used to power the neurostimulators can interfere with natural behaviors, especially when studying social interactions between multiple animals [5].

[0025] Certain embodiments include a device including a magnetic field generator and an implantable wireless neurostimulator, the implantable wireless neurostimulator including a magnetoelectric thin film; a first electrode coupled to the magnetoelectric thin film; a second electrode coupled to the magnetoelectric thin film; an electric circuit coupled to the magnetoelectric thin film; a third electrode coupled to the electric circuit; and a fourth electrode coupled to the electric circuit. In certain embodiments, the electric circuit includes a charge storage element; the magnetic field generator is configured to generate a magnetic field at an input frequency; the magnetoelectric thin film is configured to generate an electric signal when stimulated by a magnetic field at the input frequency; the electric circuit is configured to modify the electric signal to charge the charge storage element; the electric circuit is configured to modify the electric signal to transmit data to the implantable wireless neurostimulator; and the electric circuit is configured to modify the electric signal to provide a stimulation output to the third electrode and the fourth electrode.

[0026] In some embodiments, the electrical circuit is configured to modify the electrical signal to transmit data from the implanted wireless neurostimulator. In certain embodiments, the electrical circuit is formed on a complementary metal oxide semiconductor (CMOS) chip. In specific embodiments, the charge storage element is a capacitor. In certain embodiments, the electrical circuit includes a charging module, a data module, and a stimulation driver module. In certain embodiments, the charging module is configured to provide power management and charging of the charge storage element. In some embodiments, the data module is configured to provide data transmission to the implanted wireless neurostimulator. In specific embodiments, the stimulation driver module is configured to provide stimulation output to the third electrode and the fourth electrode. In certain embodiments, the charging module includes a rectifier and a reference generator. In certain embodiments, the electrical circuit is configured to modify the electrical signal to transmit data to the implanted wireless neurostimulator by performing amplitude shift modulation (ASK) of the magnetic field. In some embodiments, the magnetoelectric thin film includes a central portion, and the electrical circuit is coupled to the central portion of the magnetoelectric thin film.

[0027] In certain embodiments, the magnetoelectric film has a resonant frequency and the input frequency is within 10% of the resonant frequency of the magnetoelectric film. In certain embodiments, the magnetic field generator is configured to generate a series of magnetic field pulses. In certain embodiments, each pulse in the series of pulses has a duration of about 1 millisecond (ms). In certain embodiments, the electrical signal has an output voltage and the amplitude of the output voltage can be controlled by varying the duration of the pulses in the series of pulses. In some embodiments, the electrical signal has an output voltage and the amplitude of the output voltage can be controlled by varying the total number of pulses in the series of pulses. In specific embodiments, the electrical signal is a monophasic output signal. In certain embodiments, the electrical signal is a biphasic output signal. In certain embodiments, the magnetoelectric film includes a magnetostrictive layer and a piezoelectric layer. In some embodiments, the magnetoelectric film includes a polyvinylidene fluoride (PVDF) layer. In specific embodiments, the magnetoelectric film includes a lead zirconate titanate (PZT) layer.

[0028] In certain embodiments, the input frequency is between 20 and 500 kilohertz (kHz), and in certain embodiments, the output signal stimulates biological tissue. In some embodiments, the output signal powers an electronic device. In specific embodiments, the magnetoelectric film is a first magnetoelectric film, and the apparatus further comprises a second magnetoelectric film. In certain embodiments, the magnetic field generator comprises an electromagnet. In certain embodiments, the magnetic field generator comprises a permanent magnet. In some embodiments, the third electrode and the fourth electrode are configured as a stereotrode. In specific embodiments, the magnetoelectric film comprises a first end, a second end, a center point equidistant between the first end and the second end, and a length extending between the first end and the second end; the center portion comprises 10% of the length of the magnetoelectric film; and the center point is located at the center of the center portion of the magnetoelectric film. In certain embodiments, the magnetic field generator is configured to generate a magnetic field at an input frequency within 5% of a resonant frequency of the magnetoelectric film.

[0029] Certain embodiments include a method of generating a magnetic field by a magnetic field generator, the method including the steps of providing a device disclosed herein (e.g., including a device according to any one of claims 1-30); generating an electrical signal by a magnetoelectric thin film; and modifying the electrical signal by an electrical circuit. In certain embodiments, modifying the electrical signal by an electrical circuit includes modifying the electrical signal to charge a charge storage element; modifying the electrical signal to transmit data to an implantable wireless neurostimulator; and modifying the electrical signal to provide a stimulation output to a third electrode and a fourth electrode. Certain embodiments further include stimulating neural tissue with the modified electrical signal. Some embodiments further include powering a device with the modified electrical signal.

[0030] Certain embodiments include an implantable device including a magnetoelectric thin film; a magnetic field generator configured to generate a magnetic field at an input frequency of 20-500 kHz; and an electrical circuit coupled to the magnetoelectric thin film. In certain embodiments, the magnetoelectric thin film is configured to generate an electrical output signal when stimulated by a magnetic field at the input frequency; and the electrical circuit includes an electrical component configured to modify the electrical output signal such that the electrical output signal has an output frequency of 0.5-1000 Hz.

[0031] In specific embodiments, the magnetoelectric film includes a central portion and the electric circuit is coupled to the central portion of the magnetoelectric film. In some embodiments, the magnetoelectric film has a resonant frequency and the input frequency is within 10% of the resonant frequency of the magnetoelectric film.

[0032] In certain embodiments, the magnetic field generator is configured to generate a series of magnetic field pulses. In certain embodiments, each pulse in the series of pulses has a duration of about 1 millisecond (ms). In some embodiments, the electrical output signal has an output voltage, and the amplitude of the output voltage can be controlled by varying the duration of the pulses in the series of pulses. In specific embodiments, the electrical output signal has an output voltage, and the amplitude of the output voltage can be controlled by varying the total number of pulses in the series of pulses. In certain embodiments, the electrical output signal is a monophasic output signal. In certain embodiments, the electrical output signal is a biphasic output signal. In some embodiments, the magnetoelectric thin film comprises a magnetostrictive layer and a piezoelectric layer. In specific embodiments, the magnetoelectric thin film comprises a polyvinylidene fluoride (PVDF) layer. In certain embodiments, the magnetoelectric thin film comprises a lead zirconate titanate (PZT) layer. In some embodiments, the input frequency is between 20 and 500 kilohertz (kHz). In certain embodiments, the output signal stimulates biological tissue. In specific embodiments, the output signal powers an electronic device.

[0033] In certain embodiments, the magnetoelectric film is a first magnetoelectric film and the device further comprises a second magnetoelectric film. In certain embodiments, the magnetic field generator comprises an electromagnet. In some embodiments, the magnetic field generator comprises a permanent magnet. In specific embodiments, the electrical component comprises a full wave rectifier and a transistor. In certain embodiments, the electrical component comprises a diode and a capacitor. In certain embodiments, the device further comprises a stereotrode, and the electrical output signal is transmitted through the stereotrode. In some embodiments, the magnetoelectric film comprises a first end, a second end, a center point equidistant between the first end and the second end, and a length extending between the first end and the second end; the central portion comprises 10% of the length of the magnetoelectric film; and the central point is located at the center of the central portion of the magnetoelectric film. In specific embodiments, the magnetic field generator is configured to generate a magnetic field at an input frequency within 5% of the resonant frequency of the magnetoelectric film.

[0034] Certain embodiments include an implantable device comprising a magnetoelectric thin film; and a magnetic field generator configured to generate a magnetic field at an input frequency of 20-500 kHz. In certain embodiments, the magnetoelectric thin film is a self-rectifying thin film; the magnetoelectric thin film is configured to generate an electrical output signal at an output frequency of 0.5-1000 Hz when stimulated by a magnetic field at the input frequency. In certain embodiments, the magnetoelectric thin film comprises zinc oxide (ZnO). In some embodiments, the magnetoelectric thin film comprises a polyvinylidene fluoride (PVDF) layer doped with zinc oxide (ZnO) nanoparticles. In specific embodiments, the magnetoelectric thin film has a resonant frequency and the input frequency is within 10% of the resonant frequency of the magnetoelectric thin film. In certain embodiments, the resonant frequency of the magnetoelectric thin film is 65-75 kilohertz (kHz). In certain embodiments, the magnetic field generator is configured to generate a series of magnetic field pulses. In some embodiments, each pulse in the series of pulses has a duration of about 1 millisecond (ms). In a specific embodiment, the electrical output signal has an output voltage, and the amplitude of the output voltage can be controlled by varying the duration of the pulses in the series of pulses.

[0035] In certain embodiments, the electrical output signal has an output voltage, and the amplitude of the output voltage can be controlled by varying the total number of pulses in the series of pulses. In certain embodiments, the electrical output signal is a monophasic output signal. In some embodiments, the electrical output signal is a biphasic output signal. In specific embodiments, the magnetoelectric thin film comprises a magnetostrictive layer and a piezoelectric layer. In certain embodiments, the magnetoelectric thin film comprises a polyvinylidene fluoride (PVDF) layer. In certain embodiments, the input frequency is 20-300 kilohertz (kHz). In some embodiments, the output signal stimulates biological tissue. In specific embodiments, the output signal powers an electronic device. In certain embodiments, the magnetoelectric thin film is a first magnetoelectric thin film, and the device further comprises a second magnetoelectric thin film. In certain embodiments, the magnetic field generator comprises an electromagnet. In some embodiments, the magnetic field generator comprises a permanent magnet. In specific embodiments, the device further comprises a stereotrode, and the electrical output signal is transmitted through the stereotrode.

[0036] Certain embodiments include a method of stimulating neural tissue, comprising providing a device disclosed herein (e.g., a device according to any one of claims 1-23); generating a magnetic field with a magnetic field generator; generating an electrical output signal with an electro-magnetic film; and modifying the electrical output signal with an electrical circuit. Certain embodiments further include a method comprising stimulating neural tissue with the modified electrical output signal. Some embodiments further include powering a device with the modified electrical output signal.

[0037] Certain embodiments include a method of stimulating neural tissue, the method comprising the steps of providing an apparatus as disclosed herein (e.g., as described in any one of claims 24-43); generating a magnetic field with a magnetic field generator; and generating an electrical output signal with an electric magnetic film. Certain embodiments further include stimulating the neural tissue with the electrical output signal. Some embodiments further include powering the device with the electrical output signal.

[0038] Certain embodiments include an implantable device comprising a magnetoelectric thin film; and a magnetic field generator configured to generate a magnetic field at an input frequency of 20-500 kHz. In certain embodiments, the magnetoelectric thin film is configured to generate an electrical output signal having a voltage waveform when stimulated by a magnetic field at the input frequency; the device comprises an electrical circuit or material for modifying the voltage waveform, the electrical circuit or material comprising an electrical component configured to modify the electrical output signal such that the electrical output signal has an output frequency of 0.5-1000 Hz. In some embodiments, the material comprises zinc oxide (ZnO) nanorods.

[0039] Any embodiment of any of the methods, compositions, kits, and systems may consist of or consist essentially of (rather than include / comprise / have) the recited steps and / or features. Thus, using the words "consisting of" or "consisting essentially of" in place of any of the above open-ended linking verbs in any of the claims may modify the scope of a given claim from the scope that would result from the use of the open-ended linking verb.

[0040] Although the use of the word "or" in the claims is used to mean "and / or" unless expressly indicated to refer to alternative options only or the options are not mutually exclusive, the present disclosure supports a definition that refers only to alternatives and "and / or."

[0041] Throughout this specification, the word "about" is used to indicate that a value includes the standard deviation or error for the device or method being employed to determine the value.

[0042] In accordance with long-standing patent law, the singular indefinite articles "a" and "an," used in conjunction with the word "comprising" in a claim or specification, refer to one or more, unless otherwise noted.

[0043] [The present invention 1001] Magnetoelectric thin films; a magnetic field generator configured to generate a magnetic field at an input frequency between 20 and 500 kHz; and An electric circuit coupled to the magnetoelectric thin film Including, the magnetoelectric thin film configured to generate an electrical output signal when stimulated by the magnetic field at the input frequency; the electrical circuit includes an electrical component configured to modify the electrical output signal such that the electrical output signal has an output frequency of 0.5 to 1000 Hz; Implantable devices. [The present invention 1002] the magnetoelectric thin film includes a central portion; an electrical circuit is coupled to the central portion of the magnetoelectric thin film; The apparatus of the present invention 1001. [The present invention 1003] The magnetoelectric thin film has a resonant frequency; the input frequency is within 10% of the resonant frequency of the magnetoelectric thin film; An implantable device of the present invention 1001. [The present invention 1004] An implantable device of the present invention 1001, wherein the magnetic field generator is configured to generate a series of magnetic field pulses. [The present invention 1005] An implantable device of the present invention 1004, wherein each pulse in the series of pulses has a duration of about 1 millisecond (ms). [The present invention 1006] the electrical output signal having an output voltage; The amplitude of the output voltage can be controlled by varying the duration of the pulses in the series of pulses. An implantable device of the present invention 1004. [The present invention 1007] the electrical output signal having an output voltage; The amplitude of the output voltage can be controlled by varying the total number of pulses in the series of pulses. An implantable device of the present invention 1004. [The present invention 1008] An implantable device according to the present invention 1001, wherein the electrical output signal is a monophasic output signal. [The present invention 1009] An implantable device according to the present invention 1001, wherein the electrical output signal is a biphasic output signal. [The present invention 1010] The implantable device of the present invention 1001, wherein the magnetoelectric thin film comprises a magnetostrictive layer and a piezoelectric layer. [The present invention 1011] The implantable device of the present invention 1001, wherein the magnetoelectric thin film comprises a polyvinylidene fluoride (PVDF) layer. [The present invention 1012] The implantable device of the present invention 1001, wherein the magnetoelectric thin film comprises a lead zirconate titanate (PZT) layer. [The present invention 1013] An implantable device of the present invention 1001, wherein the input frequency is 20 to 500 kilohertz (kHz). [The present invention 1014] An implantable device of the present invention 1001, wherein the output signal stimulates biological tissue. [The present invention 1015] An implantable apparatus of the present invention 1001, wherein the output signal powers an electronic device. [The present invention 1016] The implantable device of the present invention 1001, wherein the magnetoelectric film is a first magnetoelectric film and the device further comprises a second magnetoelectric film. [The present invention 1017] An implantable device according to claim 1001, wherein the magnetic field generator comprises an electromagnet. [The present invention 1018] An implantable device according to claim 1001, wherein the magnetic field generator comprises a permanent magnet. [The present invention 1019] The implantable device of the present invention 1001, wherein the electrical components include a full-wave rectifier and a transistor. [The present invention 1020] The implantable device of the present invention 1001, wherein the electrical components include a diode and a capacitor. [The present invention 1021] An implantable device of the present invention 1001 further comprising a stereotrode, the electrical output signal being transmitted through the stereotrode. [The present invention 1022] Magnetoelectric thin film, a first end, a second end, a center point equidistant between the first end and the second end, and a length extending between the first end and the second end. Includes; the central portion comprises 10% of the length of the magnetoelectric thin film; the central point being located at the center of the central portion of the magnetoelectric thin film; An implantable device of the present invention 1002. [The present invention 1023] An implantable device according to the present invention 1001, wherein the magnetic field generator is configured to generate a magnetic field at an input frequency within 5% of the resonant frequency of the magnetoelectric thin film. [The present invention 1024] Electric and magnetic thin films; and A magnetic field generator configured to generate a magnetic field with an input frequency between 20 and 500 kHz Including, The magnetoelectric thin film is a self-rectifying thin film; the magnetoelectric thin film is configured to generate an electrical output signal at an output frequency of 0.5 to 1000 Hz when stimulated by the magnetic field at the input frequency; Implantable devices. [The present invention 1025] The implantable device of the present invention 1024, wherein the magnetoelectric thin film comprises zinc oxide (ZnO). [The present invention 1026] 1024. An implantable device according to the present invention, wherein the magnetoelectric thin film comprises a polyvinylidene fluoride (PVDF) layer doped with zinc oxide (ZnO) nanoparticles. [The present invention 1027] The magnetoelectric thin film has a resonant frequency; the input frequency is within 10% of the resonant frequency of the magnetoelectric thin film; An implantable device according to the present invention 1024. [The present invention 1028] 1027. An implantable device according to the present invention, wherein the resonant frequency of the magnetoelectric thin film is 65 to 75 kilohertz (kHz). [The present invention 1029] An implantable device of the present invention 1024, wherein the magnetic field generator is configured to generate a series of magnetic field pulses. [The present invention 1030] An implantable device of the present invention 1029, wherein each pulse in the series of pulses has a duration of about 1 millisecond (ms). [The present invention 1031] the electrical output signal having an output voltage; An implantable device of the present invention 1029, wherein the amplitude of the output voltage can be controlled by varying the duration of the pulses in the series of pulses. [The present invention 1032] the electrical output signal having an output voltage; The amplitude of the output voltage can be controlled by varying the total number of pulses in the series of pulses. An implantable device according to the present invention 1029. [The present invention 1033] An implantable device according to the present invention 1024, wherein the electrical output signal is a monophasic output signal. [The present invention 1034] The implantable device of the present invention 1024, wherein the electrical output signal is a biphasic output signal. [The present invention 1035] The implantable device of the present invention 1024, wherein the magnetoelectric thin film comprises a magnetostrictive layer and a piezoelectric layer. [The present invention 1036] 1024. An implantable device according to the present invention, wherein the magnetoelectric thin film comprises a polyvinylidene fluoride (PVDF) layer. [The present invention 1037] An implantable device of the present invention 1024, wherein the input frequency is 20 to 300 kilohertz (kHz). [The present invention 1038] An implantable device of the present invention 1024, wherein the output signal stimulates biological tissue. [The present invention 1039] The implantable apparatus of the present invention 1024, wherein the output signal powers an electronic device. [The present invention 1040] The implantable device of the present invention 1024, wherein the magnetoelectric film is a first magnetoelectric film and the device further comprises a second magnetoelectric film. [The present invention 1041] An implantable device according to claim 1024, wherein the magnetic field generator comprises an electromagnet. [The present invention 1042] An implantable device according to claim 1024, wherein the magnetic field generator comprises a permanent magnet. [The present invention 1043] An implantable device of the present invention 1024 further comprising a stereotrode, the electrical output signal being transmitted through the stereotrode. [The present invention 1044] 1. A method of stimulating neural tissue, comprising: Providing any one of the devices according to the present invention 1001 to 1023; generating a magnetic field with a magnetic field generator; generating an electrical output signal with the electromagnetic thin film; and modifying said electrical output signal with an electrical circuit. A method comprising: [The present invention 1045] The method of claim 1044, further comprising the step of stimulating neural tissue with the altered electrical output signal. [The present invention 1046] The method of any one of claims 1044 to 105 further comprising the step of powering a device with the modified electrical output signal. [The present invention 1047] 1. A method of stimulating neural tissue, comprising: Providing any one of the devices according to the present invention 1024 to 1043; generating a magnetic field with a magnetic field generator; and Producing an electrical output signal by means of an electromagnetic thin film A method comprising: [The present invention 1048] The method of claim 1047, further comprising the step of stimulating neural tissue with the electrical output signal. [The present invention 1049] The method of claim 1047 further comprising the step of powering the device with the electrical output signal. [The present invention 1050] Magnetoelectric thin films; a magnetic field generator configured to generate a magnetic field at an input frequency of 20 to 500 kHz; Including, the magnetoelectric thin film configured to generate an electrical output signal having a voltage waveform when stimulated by the magnetic field at the input frequency; the device includes an electrical circuit or material for modifying the voltage waveform; the electrical circuit or material includes an electrical component configured to modify the electrical output signal such that the electrical output signal has an output frequency of 0.5 to 1000 Hz; Implantable devices. [The present invention 1051] An implantable device according to the present invention 1050, wherein the material comprises zinc oxide (ZnO) nanorods. [The present invention 1052] A magnetic field generator; and Implantable wireless neurostimulator Including, The implantable wireless neurostimulator comprises: Magnetoelectric thin films; a first electrode coupled to the magnetoelectric thin film; a second electrode coupled to the magnetoelectric thin film; an electrical circuit coupled to the magnetoelectric thin film; a third electrode coupled to the electrical circuit; and a fourth electrode coupled to the electrical circuit; Including, the electrical circuit includes a charge storage element; the magnetic field generator configured to generate a magnetic field at an input frequency; the magnetoelectric thin film configured to generate an electrical signal when stimulated by the magnetic field at the input frequency; the electrical circuit is configured to modify the electrical signal to charge the charge storage element; the electrical circuitry is configured to modify the electrical signal to transmit data to the implantable wireless neurostimulator; the electrical circuit is configured to modify the electrical signal to provide a stimulation output to the third electrode and the fourth electrode. Device. [The present invention 1053] The device of the present invention 1052, wherein the electrical circuitry is configured to modify an electrical signal to transmit data from the implantable wireless neurostimulator. [The present invention 1054] The device of the present invention 1052, wherein the electrical circuitry is formed on a complementary metal oxide semiconductor (CMOS) chip. [The present invention 1055] The apparatus of claim 1052, wherein the charge storage element is a capacitor. [The present invention 1056] The apparatus of the present invention 1052, wherein the electrical circuit includes a charging module, a data module and a stimulus driver module. [The present invention 1057] The apparatus of the present invention 1056, wherein the charging module is configured to provide power management and charging of the charge storage element. [The present invention 1058] The apparatus of the present invention 1056, wherein the data module is configured to provide data transmission to an implantable wireless neurostimulator device. [The present invention 1059] The apparatus of the present invention 1056, wherein the stimulation driver module is configured to provide stimulation output to a third electrode and a fourth electrode. [The present invention 1060] The apparatus of claim 1056, wherein the charging module includes a rectifier and a reference generator. [The present invention 1061] The device of the present invention 1052, wherein the electrical circuit is configured to modify the electrical signal to transmit data to an implantable wireless neurostimulator data transmission by performing amplitude shift keying (ASK) of a magnetic field. [The present invention 1062] the magnetoelectric thin film includes a central portion; an electrical circuit is coupled to the central portion of the magnetoelectric thin film; The apparatus of the present invention 1052. [The present invention 1063] The magnetoelectric thin film has a resonant frequency; the input frequency is within 10% of the resonant frequency of the magnetoelectric thin film; The apparatus of the present invention 1052. [The present invention 1064] The apparatus of claim 1052, wherein the magnetic field generator is configured to generate a series of magnetic field pulses. [The present invention 1065] The apparatus of the present invention 1064, wherein each pulse in the series of pulses has a duration of about 1 millisecond (ms). [The present invention 1066] the electrical signal has an output voltage; The amplitude of the output voltage can be controlled by varying the duration of the pulses in the series of pulses. The apparatus of the present invention 1065. [The present invention 1067] the electrical signal has an output voltage; An apparatus according to the present invention 1065, wherein the amplitude of the output voltage can be controlled by varying the total number of pulses in the series of pulses. [The present invention 1068] The apparatus of claim 1052, wherein the electrical signal is a monophasic output signal. [The present invention 1069] The apparatus of claim 1052, wherein the electrical signal is a biphasic output signal. [The present invention 1070] The apparatus of claim 1052, wherein the magnetoelectric thin film comprises a magnetostrictive layer and a piezoelectric layer. [The present invention 1071] The apparatus of claim 1052, wherein the magnetoelectric thin film comprises a polyvinylidene fluoride (PVDF) layer. [The present invention 1072] The apparatus of claim 1052, wherein the magnetoelectric thin film comprises a lead zirconate titanate (PZT) layer. [The present invention 1073] The apparatus of the present invention 1052, wherein the input frequency is 20 to 500 kilohertz (kHz). [The present invention 1074] The apparatus of claim 1052, wherein the output signal stimulates biological tissue. [The present invention 1075] The apparatus of the present invention 1052, wherein the output signal powers an electronic device. [The present invention 1076] The apparatus of claim 1052, wherein the magnetoelectric film is a first magnetoelectric film and the apparatus further comprises a second magnetoelectric film. [The present invention 1077] The apparatus of claim 1052, wherein the magnetic field generator comprises an electromagnet. [The present invention 1078] The apparatus of claim 1052, wherein the magnetic field generator comprises a permanent magnet. [The present invention 1079] The apparatus of claim 1052, wherein the third electrode and the fourth electrode are configured as a stereotrode. [The present invention 1080] Magnetoelectric thin film, a first end, a second end, a center point equidistant between the first end and the second end, and a length extending between the first end and the second end. Includes; the central portion comprises 10% of the length of the magnetoelectric thin film; the central point being located at the center of the central portion of the magnetoelectric thin film; The apparatus of the present invention 1052. [The present invention 1081] The apparatus of claim 1052, wherein the magnetic field generator is configured to generate a magnetic field at an input frequency within 5% of the resonant frequency of the magnetoelectric thin film. [The present invention 1082] 1. A method of generating a magnetic field by a magnetic field generator, comprising: Providing any one of the devices according to the present invention 1052 to 1081; generating an electrical signal by the magnetoelectric thin film; and The process of modifying an electrical signal using an electrical circuit A method comprising: [The present invention 1083] modifying the electrical signal by an electrical circuit; modifying the electrical signal to charge a charge storage element; modifying the electrical signal to transmit data to an implantable wireless neurostimulator; and modifying the electrical signal to provide a stimulation output to a third electrode and a fourth electrode. The method of the present invention 1082, comprising: [The present invention 1084] The method of claim 1082, further comprising the step of stimulating neural tissue with the altered electrical signal. [The present invention 1085] The method of claim 1082 further comprising the step of powering the device with the modified electrical signal. Other objects, features and advantages of the present invention will become apparent from the following detailed description. It should be understood, however, that the detailed description and specific examples, while indicating particular embodiments of the invention, are given by way of illustration only, since various changes and modifications within the spirit and scope of the invention will become apparent to those skilled in the art from this detailed description.

[0044] The following drawings form part of the present specification and are included to further illustrate certain aspects of the present invention. The invention may be better understood by reference to one or more of these drawings in combination with the detailed description of specific embodiments presented herein. The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication containing color drawing(s) will be provided by the Office upon request and payment of the necessary fee. [Brief description of the drawings]

[0045] [Figure 1] FIG. 1 is a schematic diagram of an apparatus according to an exemplary embodiment of the present disclosure. [Diagram 2] 1 illustrates various waveforms of an exemplary embodiment of the present disclosure. [Diagram 3] 1 illustrates various circuits that can be used to generate various waveforms of the present disclosure. [Figure 4] 1 illustrates various aspects of fixed amplitude and variable amplitude configurations. [Diagram 5] FIG. 1 is a schematic diagram of a system according to an exemplary embodiment of the present disclosure. [Figure 6] We present aspects of how magnetoelectric (ME) thin films convert alternating magnetic fields into voltages. [Figure 7] The relationship between film properties and film length is shown. [Figure 8]The thin film output voltage is shown as a function of the bias magnetic field. [Figure 9] 1 illustrates aspects of a magnetic field driver. [Figure 10] 1 shows aspects of a monophasic ME stimulator for activating cells in vitro. [Figure 11] 1 shows the aspects of cells grown directly on ME thin films. [Figure 12] 1 shows aspects of a biphasic ME stimulator that activates neurons in ex vivo brain slices. [Figure 13] 1 shows aspects of a biphasic stimulator for deep brain stimulation (DBS) in freely moving rats. [Figure 14] A schematic diagram showing the position and spacing of the wires and the number of drivers used to generate the alternating magnetic field in the experimental setup is shown. [Figure 15] A COMSOL simulation of the magnetic field on a circular coil and the measured device output voltage as a function of distance on the coil are shown. [Figure 16] 1 shows aspects of miniaturized multi-channel stimulation in a human skull model. [Figure 17] 1 shows voltage versus time for PVDF / Metglas thin films doped with ZnO nanoparticles. [Figure 18] 1 shows X-ray diffraction data illustrating the poling treatment of PVDF along with the α and β phase peaks of PVDF. [Figure 19] 1 illustrates aspects of ZnO / PVDF thin film fabrication. [Figure 20] 13 shows the self-rectification of electrical current in a PVDF / ZnO nanowire hybrid material used as an energy harvester when a finger presses on the material. [Figure 21] FIG. 1 shows a circuit diagram of an exemplary embodiment of the present disclosure. [Figure 22A] FIG. 1 shows an electrical circuit diagram of an exemplary embodiment of the present disclosure. [Figure 22B] Schematics of the active rectifier, native MOS-based voltage reference and 5-bit programmable stimulus driver are shown. [Diagram 23]1 illustrates charge, data and stimulus waveforms according to an exemplary embodiment of the present disclosure. [Figure 24] The characteristics of the ME transducer are shown. (a) Equivalent circuit model of the ME material and preliminary COMSOL simulation. (b) The induced voltage is maximized at the mechanical resonance. (c) Impedance of the ME transducer (PZT based) versus magnetic field frequency. [Diagram 25] The available power under impedance matching between the ASIC and the ME transducer is shown for (a) purely resistive impedance and (b) capacitive impedance cases. [Figure 26] 1 shows the proposed adaptive data reception scheme: (a) circuit for mode switching; and (b) adaptive threshold voltage generation and data reception. [Figure 27] 4 shows simulation waveforms of adaptive data reception. [Figure 28] Low-power CMOS circuit components: (a-b) calibration-free voltage / current reference, (c) measurement results including manufacturing variations, (d) programmable stimulus driver, and (e) Wheatstone bridge-based timing reference. [Figure 29] 1A-1D show exploded and assembled views of an exemplary embodiment of the present disclosure. [Diagram 30] A table of encapsulation strategies is shown. [Diagram 31] A wearable, battery-powered magnetic field transmitter is shown. (A-C) Shows the concept of a wearable, battery-powered magnetic field generator capable of driving an implanted ME neurostimulator. (D) Shows a preliminary desktop unit demonstrating a 1 mT magnetic field generated by a 3.7 mA current passing through the coil, drawing only 0.2 mA on the battery. [Diagram 32]In vitro testing is shown. (A) Immersion testing in PBS at 37°C (or 80°C) can be used to confirm stable hermetic packaging. Preliminary data shows 28-day immersion testing of polyimide encapsulated magnetoelectric thin films showing stable peak-to-peak voltage. (B) To measure the fully packaged CMOS devices, a flexible PCB can be used to record the voltage waveform in a wired configuration or the LED intensity and frequency can be monitored in a wireless configuration. Alternatively, a fiber can be coupled to the LED (not shown) if unstable measurements are seen by the camera. [Diagram 33] Principles and characteristics of magnetoelectric (ME) transducers; comparison of wireless power transmission mechanisms; diagrams of the proposed MagMote and wearable neurostimulation system concepts are shown. [Diagram 34] The system diagram of the adaptive operation scheme; robustness improvement of the four-input comparator in tracking Vrect; robustness analysis of the clock; measured data modulation depth; and Istim,max under varying Vs are shown. [Diagram 35] MagMote waveforms measured at various stimulation frequencies; zoomed-in views of biphasic stimulation showing power recovery, data transmission and amplitudes of 0.05-1.5 mA, pulse widths of 256 and 512 µs are shown. [Diagram 36] In vitro setup in PBS (top left); ME power delivery at various distances and stimulus waveforms (top right); Vrect and Pin,avg measured in various media and the stability of Pin,avg over a 7 day immersion test (bottom). [Figure 37] The experimental setup for the Hydra test (top left); time-aligned measured fluorescence increase in response to electrical stimulation (top right); and a comparison with a state-of-the-art wireless implantable neurostimulator (bottom). [Figure 38] Chip micrographs, SoC performance overview and power breakdown are shown. [Figure 39]The experimental setup to capture fluorescence images using a microscope and constant excitation light; the stereotrode connected to the proposed device; and the laser used to image GCaMP6 activity in Hydra are shown. [Diagram 40] 1 shows a portable magnetic field driver and its system diagram. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0046] Detailed Description of the Invention Referring first to Figure 1, a basic schematic diagram illustrates aspects of an apparatus 10 of the present disclosure. As described in more detail below, exemplary embodiments of the present disclosure can be configured to stimulate neural tissue, power devices that record physiological signals, and / or combine recording and stimulation for closed-loop bioelectronic medicine.

[0047] The device 10 includes a magnetic field generator 20 (shown in this example as a coil), a magnetoelectric (ME) film 30, and an electrical circuit 40 coupled to a central portion of the film 30. Coupling of the circuit to the center of the film is important for maximum efficiency as it minimizes mechanical damping of the ME film resonance, allowing for a voltage amplitude capable of stimulating cells.

[0048] It will be understood that the schematic diagram shown in Figure 1 is a simplified representation of an embodiment disclosed herein, which is described in more detail below. In general, a magnetic field generator 20 generates a magnetic field, which is converted by a membrane 30 into a high frequency monophasic or biphasic voltage / current, which is then converted by an electrical circuit 40 into a conditioned stimulus.

[0049] The device 10 can be configured to control the amplitude and waveform of the conditioned stimulation to achieve stable and effective neural stimulation. In particular, the electrical circuit 40 can be configured to convert the high frequency magnetic field into any of the following waveforms shown in Figure 2. Figure 2 shows a monophasic high frequency waveform (e.g., created using a diode), a monophasic low frequency / envelope waveform (e.g., created using a diode and a capacitor), a biphasic high frequency waveform (e.g., created using a circuit and two membranes) and a biphasic low frequency / envelope waveform (created using a circuit and two membranes or a circuit and one membrane).

[0050] Referring now to FIG. 3, examples of various circuits that can be used to generate various waveforms in accordance with the present disclosure. It will be understood that the diagram shown in FIG. 3 is merely a schematic diagram and is not to scale. For example, in certain examples, the circuit shown in FIG. 3 is coupled to an end of the membrane, however, exemplary embodiments include circuitry coupled to a central portion of the membrane. It will also be understood that the diagram shown is merely illustrative of the type of circuitry that can be utilized to generate a desired waveform. As shown in FIG. 3, the exemplary electrical circuit can include rectifiers, diodes, transistors, resistors, and other components as necessary to generate a desired waveform.

[0051] In certain embodiments, the piezoelectric thin film acts as a capacitor with a capacitance value determined by its size and dielectric constant. Adding a diode in parallel with the thin film creates a clamper circuit that biases the high frequency waveform. In a two-thin film configuration, a transistor can be added so that the circuit leading to the inactive thin film is switched off while the active thin film provides a signal to enable the creation of a biphasic waveform. Using a resistor can limit the current or voltage of the waveform applied to the tissue and provide a place to probe the circuit to test functionality. Using an LED or Zener diode in parallel with the thin film and diode can cap the thin film output voltage with the forward bias voltage of the Zener diode or LED. Using a capacitor in parallel with the thin film can smooth the waveform and create the envelope shape shown in FIG. 2. In addition to using discrete components, an integrated circuit can be designed to generate the desired waveform, which can also include a capacitor for energy storage.

[0052] The need to create monophasic or biphasic stimulation depends on which condition is being treated. For example, chronic back pain may be addressed by monophasic stimulation, while Parkinson's disease may be addressed by biphasic stimulation, as shown in panel D of FIG. 6.

[0053] To maintain stable efficacy of the device, it is often important to stabilize the voltage output as the device moves relative to the magnetic field generator (e.g., fixed amplitude devices). It may also be important to adjust the amplitude of stimulation as tissue changes, as the device moves over time, or as the patient's needs change (variable amplitude).

[0054] Now referring to FIG. 4, various aspects of fixed amplitude and variable amplitude configurations are shown. To condition the amplitude of stimulation, the device can be operated in fixed amplitude configuration, which can be achieved by adding an LED or diode in parallel with the first diode. In such an embodiment, the forward bias voltage sets the upper voltage limit of the device. These configurations can be useful for applications where the magnetic field strength varies over a wide area.

[0055] The device can also be operated in a variable amplitude configuration to condition the amplitude of the stimulation. Such a configuration can use pulse width modulation to increase or decrease the stimulation amplitude by varying the duration or frequency of the high frequency magnetic field pulses.

[0056] Certain embodiments may utilize frequency shifting. For example, the output voltage can be lowered by shifting the magnetic field frequency slightly away from the resonance of the characterized thin film. This is useful for applications where the pulse length needs to be increased and the amplitude needs to be kept low.

[0057] Certain embodiments may also use pulse duration control. For example, for very short pulses, the thin film can be characterized and the calculated rise time constant can be used to determine the number of magnetic field pulses required to achieve the desired voltage. This can be useful for amplitude control in applications where the pulse time is very short.

[0058] Additionally, certain embodiments may utilize programmable functions to program the device, for example transmitting data via a magnetic field to switch stimulation modes in an electrical circuit as needed.

[0059] Although the resonant frequency of the ME thin film is too fast for cells to respond (~50-500 kHz), we can apply a slowly varying envelope to this high-frequency carrier to stimulate electrically excitable cells. For example, we can create a stimulation pulse train at 130 Hz (a typical therapeutic frequency for DBS) by applying a series of short high-frequency pulses (~1 ms duration) and rectifying the ME response using a diode.

[0060] Exemplary embodiments can dynamically adjust the voltage generated by the ME device to match conventional DBS stimulation by changing the nature of the magnetic stimulation. For example, by adding a capacitor and resistor in series with the ME membrane, an RC time constant can be manipulated that is faster than the desired pulse train (130 Hz) and slower than the carrier frequency (approximately 100 kHz). In this case, the duration or duty cycle of the high frequency magnetic pulse can be adjusted so that the membrane reaches a voltage determined by the time constant for charging the membrane.

[0061] Referring now to FIG. 5, a more detailed schematic diagram of a system 100 of an exemplary embodiment of the present disclosure is provided. The system 100 includes a first magnetoelectric film 110 and a second magnetoelectric film 120. It will be understood that other embodiments may include only one magnetoelectric film. The system 100 further includes a magnetic field generator 130 including a permanent magnet 132 and an alternating magnetic coil 134. The illustrated embodiment utilizes a permanent magnet in the magnetic field generator, although other embodiments may utilize an electromagnet in the magnetic field generator. In certain embodiments, the magnetic field generator 130 is configured to generate a magnetic field 136 at a frequency between 20-500 kHz.

[0062] In the illustrated embodiment, the first and second magnetoelectric films 110 and 120 are configured to generate an electrical output signal 146 when stimulated by a magnetic field 136 at an input frequency generated by a magnetic field generator 130. In certain embodiments, the electrical circuit 140 is configured to modify the electrical output signal 146 such that the electrical output signal 146 has an output frequency between 0.5 and 1000 Hz. As described in further detail below, the electrical circuit 140 may include various components, such as one or more rectifiers 142, transistors 144, and resistors 148. The electrical output signal 146 may be transmitted through a stereotrode 150 to activate neurons in a biological sample 160 (e.g., brain tissue).

[0063] In certain exemplary embodiments, the magnetoelectric films 110 and 120 can be fabricated by combining a piezoelectric layer, a polyvinylidene fluoride (PVDF), and a magnetostrictive layer (e.g., Metglas) to form a magnetoelectric (ME) film. In a specific embodiment, the two layers are first plasma treated to promote adhesion and a clean surface. The layers can then be bonded together by application of an epoxy layer about 30 μm thick. Once cured, the composite film can be cut using a femtosecond laser system capable of cutting devices at the sub-millimeter scale.

[0064] To interrogate the high voltage on the membrane, the microwires can be bonded to the center portion of the membrane using silver printing or silver conductive epoxy. The electrical circuit 140 can include a novel H-bridge circuit to act as a magnetic coil driver to reach high voltages and frequencies. It will be understood that the specific fabrication methods and materials disclosed herein are merely exemplary, and that aspects of the present disclosure are not limited to such methods and materials.

[0065] In certain embodiments, the magnetoelectric thin films 110 and 120 can be fabricated by spin-coating a piezoelectric polymer (PVDF-TrFE), using PZT piezoelectric ceramic material, or using electrospinning and direct writing of PVDF nanofibers. In a specific embodiment, when the ME thin film comes into contact with an alternating magnetic field generated by a driver and a permanent bias magnet or a DC bias coil, the film resonates and generates up to 50V peak-to-peak. Appropriate circuitry can be electrically coupled to the thin film to obtain a desired output voltage waveform, examples of which are described below.

[0066] In certain embodiments, other magnetostrictive materials may be utilized, including PVDF thin films doped with zinc oxide (ZnO) nanoparticles. Having a high piezoelectric coefficient, such as lead zirconate titanate (PZT), is beneficial in the fabrication of materials for higher energy output and harvesting. PVDF is a piezoelectric polymer with lower piezoelectricity than PZT, but it is a soft conformal material, and it has been shown that ZnO nanoparticles embedded in PVDF increase its piezoelectricity, primarily by increasing the dielectric constant of the composite thin film

[48] .

[0067] A detailed description of the embodiment incorporating ZnO nanoparticles and further explanation of the operating principle can also be found in the Examples and Results section below. EXAMPLES

[0068] The following examples are included to demonstrate preferred embodiments of the present disclosure. It should be understood by those skilled in the art that the techniques disclosed in the following examples represent techniques that the inventors have found to work well in implementing the present disclosure, and therefore can be considered to constitute preferred modes for its implementation. However, it should be understood by those skilled in the art in light of the present disclosure that numerous changes can be made to the specific embodiments disclosed without departing from the spirit and scope of the disclosure, and still obtain the same or similar results.

[0069] design challenges One of the main challenges for wireless neurostimulators is to create efficient miniature devices (<1 cm in length) that operate reliably beneath bone and tissue as the animal or human patient engages in normal activities. At lengths less than 1 cm, devices can be fully implanted in the periphery and light enough to allow unrestricted animal behavior, but power delivery remains a challenge for devices this small. Efficient power transmission by propagating electromagnetic waves requires antennas with feature sizes comparable to the electromagnetic wavelength. Thus, for sub-millimeter devices such as the proposed RF-powered "neuro-brain" [6], the effective power transmission frequency is in the GHz range, where electromagnetic radiation is absorbed by the body [7]. This absorption of radio-frequency electromagnetic energy limits the amount of power that can be safely delivered to implants deep within tissue [7]. As a result, researchers typically turn to magnetic induction or batteries to power implanted devices. However, these technologies also limit the degree of miniaturization. Batteries increase the size of the device and add significant weight. In addition, batteries need to be replaced or recharged, which may limit potential applications. On the other hand, inductively coupled coils can be made smaller and lighter than batteries, but the power a receiver coil can generate is directly related to the amount of magnetic flux captured by the area of ​​the coil. Thus, as receiver coils are made smaller, they have less output power and are more sensitive to perturbations in the distance or angle between the transmitter and receiver [8]. For example, Freeman et al. demonstrated that a small inductive coil, less than 1 mm in diameter, can power a stimulator for the sciatic nerve in anesthetized rats [9]. However, in its current form, this device has difficulty providing stable stimulation in freely moving animals due to the decrease in power coupling efficiency with changes in angle and distance between the receiver and transmitter coils.

[0070] In addition, for neurostimulators to treat many neurological disorders, such as Parkinson's disease (PD), obsessive-compulsive disorder, and epilepsy, they must operate safely and effectively in the high-frequency "therapy band" of 100-200 Hz [10-12]. This type of high-frequency neurostimulation is challenging because the charge on the electrodes must be dissipated between successive stimulation pulses to prevent electrolysis, tissue damage, and changes in local pH

[13] . Charge dissipation at high frequencies is achieved by using biphasic stimulation waveforms that actively or passively charge and discharge the electrodes with each cycle. In fact, all clinically approved electrical neurostimulation therapies in this therapy band use various forms of "charge-balanced" biphasic stimulation waveforms

[14] .

[0071] Recently, several promising alternatives to magnetic induction and batteries have enabled miniaturized neurostimulators. However, these techniques have yet to demonstrate in vivo operation in therapeutic high frequency bands in freely moving animals. Montgomery et al. and Ho et al. have shown that the mouse body can be used as an electromagnetic resonant cavity to effectively transmit energy to subwavelength-scale devices implanted inside the animal's body [15, 16]. This technique is particularly effective for driving small LEDs for optogenetic stimulation. However, electrical stimulation is limited to <20 Hz because the electrical waveform is monophasic. By absorbing energy from a high frequency (500 kHz) magnetic field using superparamagnetic nanoparticles

[17] , specific regions of the brain in freely moving animals can be heated [18, 19]. This localized heating can stimulate neural activity if the targeted brain region is genetically modified to respond to temperature changes [18, 19]. However, this technique has not yet demonstrated high frequency operation because it requires genetic modification (adding regulatory complexity) and also requires cooling of the tissue between stimulation intervals. Ultrasound offers a promising and efficient method for powering bioelectronic implants because ultrasonic wavelengths are 105 times smaller than electromagnetic waves at the same frequency, allowing sub-millimeter-sized devices to have wavelength-scale piezoelectric antennas [20, 21]. However, the realization of these “neural dust” motes can be challenging in freely moving animals because the impedance mismatch between air, bone and tissue requires contact between the soft tissue and the ultrasound transducer for efficient power transmission. As a result, ultrasound-powered neurostimulators in freely moving animals have yet to be demonstrated

[22] .

[0072] Here, we show that magnetoelectric (ME) materials enable the first magnetically powered miniature neurostimulators that operate in the therapeutic radio frequency band. Similar to induction coils, these materials convert a magnetic field into an electric field, but instead of using an embedded coil, we use materials that generate voltage through mechanical coupling between a magnetostrictive layer and a piezoelectric layer in a thin film. That is, a magnetic field generates a strain in the magnetostrictive layer when the magnetic dipole aligns with the applied magnetic field. That strain exerts a force on the piezoelectric layer, which generates a voltage (Figure 5). By utilizing this transduction mechanism, magnetoelectric materials are not subject to the same miniaturization constraints as coils and can be driven by weak magnetic fields on the order of a few millitesla. These properties have led researchers to propose magnetoelectric materials as promising materials for bioelectronic implants [23-27]. Here, we demonstrate the first proof-of-principle wireless neurostimulator based on ME materials in a freely moving rodent model for Parkinson's disease (PD) and demonstrate that these materials can power miniature devices deep within the human brain.

[0073] Fabrication and characterization of ME stimulator We fabricated a proof-of-principle ME stimulator by bonding a rectangular magnetostrictive layer (Metglas) to a platinum-coated piezoelectric layer polyvinylidene fluoride (PVDF). We then encapsulated the thin film in a protective Parylene C layer (8-10 μm thick) (see panel (a) of FIG. 6; also see the Methods section). This embodiment incorporated a PVDF layer of 28-110 μm, thereby yielding a total device thickness of 50-150 μm. Other embodiments may include alternative or additional piezoelectric materials, including, for example, lead zirconate titanate, also known as PZT.

[0074] FIG. 6 shows an overview of how the ME thin film converts an alternating magnetic field into a voltage. In panel (a) of FIG. 6, a diagram of the ME device in a freely moving rat for wireless nerve stimulation is shown. In the illustrated embodiment, the active ME element includes a piezoelectric PVDF (blue) and Metglas (grey) laminate encapsulated by Parylene C. The inset shows the working principle that the strain generated when magnetizing the grey magnetostrictive layer is transferred to the blue piezoelectric layer, which creates a voltage across the thin film. In panel (b) of FIG. 6, an example of a resonance response curve of the ME thin film shows that the maximum voltage is generated when the magnetic field frequency matches the acoustic resonance of 171 kHz. Measuring the voltage across the thin film showed a dramatic voltage increase when the applied magnetic field frequency matches the acoustic resonance frequency. The photo inset shows an example of an assembled ME stimulator. The "Stress Profile" inset shows a top view of the stress generated in the ME thin film, calculated by finite element simulation on / off resonance (COMSOL). Panel (c) of Figure 6 shows a device testing apparatus equipped with a permanent magnet for applying a bias magnetic field and an electromagnetic coil for applying an alternating magnetic field (scale bar: upper figure = 1 cm, lower figure = 2 mm).

[0075] Panel (d) of Figure 6 shows the maximum stimulation duration of the ME device in biphasic and monophasic operation (using repeated 400 μs / phase pulses with increasing frequency). In this example, the maximum stimulation time is determined by the time of electrolysis on a stereotrode in saline, evidenced by air bubbles (error bars ±1 standard deviation, n=4 trials). The dashed lines indicate the frequencies of electrical stimulation used in various clinical applications, indicating that biphasic operation is necessary for many clinically relevant applications. Roman numerals indicate stimulation frequencies demonstrated by previously published miniature magnetic stimulators (i: Magnetothermal, Chen et. al, 2015; ii: Magnetothermal, Munshi et. al, 2017; iii: Mid-Field Optogenetics, Montgomery et. al, 2015; iv: RF Inductive Coupling, Freeman et. al, 2017).

[0076] Since the resonant frequency is proportional to the inverse of the membrane length, multiple membranes can be designed and selectively activated by varying the stimulation frequency, as shown in FIG. 7. Using this principle, we can use different magnetic field frequencies to activate separate devices that may be in different regions of the body, or create a biphasic stimulator by interleaved resonant stimulation of two different membranes, each driving a positive or negative phase of neural stimulation. FIG. 7 shows the ME characteristics as a function of membrane size. Panel (a) shows a schematic of the experimental setup used to collect the data. Tests were performed on ME membranes of three different PVDF thicknesses: 28 (blue), 52 (red) and 110 (yellow) μm. Panel (b) of FIG. 7 shows the resonant frequency as a function of membrane length, panel (c) shows the output voltage as a function of membrane length, panel (d) shows the output voltage as a function of membrane surface area, and panel (e) shows the Q value as a function of membrane length.

[0077] Panel (f) of Figure 7 shows the maximum output as a function of membrane width for a 52 um thick PVDF, and panel (g) shows the peak resonant voltage plotted against PVDF thickness, indicating that the peak ME voltage increases with PVDF thickness. Error bars indicate ±1 standard deviation for n≈50 membranes per thickness. Panel (h) of Figure 7 shows the ME voltage as a function of the angle between the membrane and the coil. The blue region indicates the range of actuation angles where the voltage is greater than the expected stimulation voltage.

[0078] The voltage generated by the ME thin film can be further enhanced by applying a constant bias field by a permanent magnet or electromagnet, as shown in panel (c) of Figure 6. Since the strain of magnetostrictive materials is a sigmoidal function of the magnetic field strength, the change in voltage generated by an alternating magnetic field is greatest when the magnetic field oscillates around the midpoint of the sigmoid, as shown in Figure 8 [28, 29]. Figure 8 shows the thin film output voltage as a function of the bias field. The peak resonant voltage is significantly enhanced by a moderate bias field, which can be generated by a permanent magnet.

[0079] Thus, a bias field can be used to offset the magnetic field near the center of the sigmoidal magnetostrictive response curve. This bias field allows us to generate therapeutic voltage levels while applying a small (a few mT) alternating magnetic field using an electromagnetic coil and a custom control circuit that specifies the frequency and timing of the alternating magnetic field, as shown in FIG. 9. FIG. 9 shows aspects of the magnetic field driver. Panel (a) of FIG. 9 shows a schematic diagram of the main components of the magnetic field driver, with dashed lines indicating the components depicted in panel (b). Panel (c) of FIG. 9 shows the output waveform for monophasic stimulation and the parameters that can be controlled by the driver software, and panel (d) shows the output waveform for biphasic stimulation and the parameters that can be controlled by the driver software.

[0080] To identify the operating frequency of the ME stimulators, we tested them in saline and found that in the case of biphasic stimulation waveforms, we could apply constant stimulation up to at least 800 Hz without significant hydrolysis. For this test, we actuated either one membrane for monophasic stimulation or two membranes for biphasic stimulation mounted on a stereotrode (microprobe) in saline (see "Methods"). We then measured the time at which we could notice an air bubble at the electrode tip resulting from hydrolysis. This hydrolysis event is indicative of a condition that would damage the surrounding tissue. We found that in the case of monophasic stimulation waveforms, stimulation frequencies above 50 Hz produce hydrolysis, while biphasic charge-balanced stimulation shows no hydrolysis up to the maximum tested frequency of 800 Hz. Compared to previously demonstrated miniature magnetic nerve stimulators, the biphasic ME device presented here is the first to access the high frequency bands used in clinical applications such as the treatment of Parkinson's disease and obsessive-compulsive disorder (shown in panel (d) of Figure 6).

[0081] An additional challenge of any wirelessly powered neurostimulator is maintaining a well-controlled stimulation voltage. This challenge becomes especially prevalent as devices become smaller, which often reduces power transfer efficiency and creates increased sensitivity to alignment between the device and the power transmitter. ME materials offer two main advantages that can enable stable and effective stimulation even as devices become smaller and move relative to the driver coil.

[0082] First, the ME device generates voltages well in excess of the effective stimulation potential, allowing it to be effective even when the material is not aligned with the driver coil. At resonance, we measured ME voltages of over 30V at a field strength of only 1mT (shown in panels (c) and (d) of Figure 7). Because the effective stimulation voltage is typically 1-5V, we can use an LED or Zener diode to cap the applied voltage to this effective stimulation range. As long as the voltage generated by the ME thin film is equal to or greater than the capped voltage, we apply approximately the same stimulation voltage regardless of the angle or distance between the driver coil and the ME thin film. We found that for a typical thin film, we can turn the film ±80° and maintain a voltage of over 3V (shown in panel [h] of Figure 7). This large angular tolerance is aided by the high permeability of the Metglas layer, which helps to orient the magnetic field lines along the long axis of the thin film (where it is most effective in creating a magnetostrictive response).

[0083] Second, the voltage generated by the piezoelectric material depends on the thickness of the piezoelectric layer, not on the area of ​​the film

[30] , allowing us to fabricate small magnetoelectric films that generate nearly the same stimulation voltage as larger devices. Figure 7 shows the peak voltage and Q value generated for ME films of various areas. We found that for a 52 μm thick PVDF layer, the voltage remains around 10 V, even as the film length is reduced. The ±40% variation in peak voltage and Q value is likely due to defects introduced during film fabrication and could be reduced by improvements in fabrication. We also verified that the output voltage depends only on the piezoelectric film thickness by measuring the peak voltage from ME devices with three different thicknesses of PVDF (28 μm, 52 μm, and 110 μm). As expected, it can be seen that the peak voltage increases linearly with PVDF thickness and is independent of the film thickness. We calculate (and experimentally confirm) that for a given thickness and length / width ratio >3, the power generated by an ME device is proportional to the membrane width (see panel [f] in Figure 7). We calculate that a membrane less than 1 cm long can generate up to 4 mW, even though the power decreases as the membrane gets smaller. This is more than enough for many wireless applications, including neural stimulation

[31] .

[0084] Monophasic stimulation with ME thin films evokes action potentials in vitro Using fluorescence microscopy to image voltages in cultured cells, we found that monophasic stimulation with ME thin films at 100 Hz for 50 ms reliably stimulated action potentials (APs). For these experiments, we used a "spiking" human embryonic kidney (HEK) cell line modified to express sodium and potassium channels. These cells have a spike-like electrical waveform that is rectangular, which can last for several seconds depending on the confluency of the culture

[32] . To determine the relative timing between magnetic stimulation and action potential generation, we transfected these cells with ArcLight33, a genetically encoded voltage indicator that allows us to measure action potentials using fluorescence microscopy.

[0085] Referring now to FIG. 10, aspects of a monophasic ME stimulator for activating cells in vitro are shown. Panel (a) of FIG. 10 shows a schematic of the experimental setup, and panel (b) shows a microscopic image of holes punched in the ME thin film, with finite element simulation of the electric field showing that the holes generate fringing electric fields that overlap with the cultured cells. Panels (c) and (d) of FIG. 10 show the voltage across the ME thin film when the magnetic field is on-resonance and off-resonance, respectively. The inset shows a close-up of the high-frequency carrier waveform. Panels (e-g) of FIG. 10 show fluorescence from a spiking HEK transfected with ArcLight, showing that action potentials are induced by the ME thin film driven on-resonance (panel [e]), but not when the film is driven off-resonance (panel [f]). Fluorescence from GFP-transfected HEK cells showed no response when the ME thin film was driven in the on-resonance state, confirming that the measured ArcLight responses were the result of changes in membrane potential difference and not an artifact of magnetic field or acoustic resonance of the ME thin film.

[0086] To image the fluorescence upon application of a magnetic field, we developed an experimental setup that allowed the placement of cells and ME thin films under an objective lens at the center of a 10 cm long solenoid (with a 3 cm gap). This configuration allowed the ME thin film, cells, and objective lens to be placed at the center of the applied magnetic field, as shown in panel (a) of Figure 10. Two slightly larger coils placed on either side of the gap provide a constant bias field.

[0087] We then approximated an embedded ME stimulator using two experimental setups: (1) growing cells directly on the ME thin film (shown in FIG. 11); and (2) placing a coverslip with cells attached on top of the ME thin film (shown in FIG. 10). FIG. 11 shows ME stimulation of cells grown directly on the ME thin film. Panel (a) of FIG. 11 shows a schematic of the experimental setup. Panel (b) of FIG. 11 shows a microscopic image of fixed cells attached to the area around the punched holes (Hoechst / Calcein-AM, labeling cells before fixation). Panel (c) of FIG. 11 shows ArcLight fluorescence of spiking HEK cells when the magnetic field is on-resonance, and panel (d) of FIG. 11 shows that when the magnetic field is off-resonance.

[0088] To culture cells directly on the ME thin film, we coated the top parylene layer with poly-l-lysine. Healthy growth of HEK on the ME device indicates that this encapsulation method prevents the ME material from restricting cell growth (Figure 11, panel [b]). However, in a typical use case, target cells may not attach to the ME stimulator, so we also tested the response of cells placed on the ME material. In this configuration, we first grew cells on the coverslip for 3-5 days, then inverted the coverslip and placed it on the ME for testing (Figure 6, see "Methods").

[0089] To create fringe electric fields that interact with cultured cells, we punched holes in the ME thin film, as shown in panel (b) of Figure 10. The thin film was otherwise fabricated as described herein (Figure 6, "Methods"). In experiments using ME thin films and Pt electrodes, we found that high-frequency biphasic stimulation at ME resonant frequencies (typically 20-150 kHz) was ineffective at stimulating action potentials (APs) in cultured HEK, as predicted by the low-pass filtering properties of the cell membrane [9]. To create an effective monophasic stimulation waveform, we rectified the voltage using a Schottky diode to create a fully positive or negative voltage waveform, depending on the diode orientation. This rectified waveform has a monophasic envelope that varies slowly within the <500 Hz frequency band to which cells respond (panels [c] and [d] of Figure 10).

[0090] For both cells grown directly on ME thin films and cells placed in contact, we found that five stimulation pulses with an envelope frequency of 100 Hz consistently stimulated APs in spiking HEK cells (panel [e] in Figure 10, panel [c] in Figure 11). Significantly, this 10-500 Hz stimulation frequency spans the therapeutic window of many deep brain stimulation therapies

[34] and is difficult to achieve with other wireless stimulators that compensate for low-efficiency energy harvesting by charging on-board capacitors

[35] . For the experiments disclosed herein, the carrier frequency of the magnetic field was the resonant frequency of the device, which varied between 20-40 KHz depending on the device length. To test stimulation reliability, we repeated the five-pulse stimulation three times over a 30-second period. We observed APs for every stimulation pulse in n=43 cells on coverslips and n=144 cells grown on thin films. By imaging voltage-sensitive fluorescence when the magnetic field was tuned out of resonance, we confirmed that the AP stimulated by the ME thin film was indeed the result of resonant excitation of the film and not an artifact of the applied magnetic field. In the case of non-resonant excitation, we found no correlation between the applied magnetic field and the fluorescently detected AP in spiking HEKs (panel [f] in Fig. 10, panel [d] in Fig. 11), supporting the conclusion that the AP was stimulated by the ME thin film in a resonant state. We also confirmed that the recorded fluorescent signal indeed represented a voltage-dependent ArcLight response by imaging cells transfected with a voltage-independent cytoplasmic GFP. These cells showed no change in fluorescence when the thin film was driven at the resonant frequency (panel [g] in Fig. 10).

[0091] Biphasic stimulation with ME thin films evokes action potentials in brain slices. As mentioned above, biphasic stimulation is preferred for most applications due to the desire to create a charge-balanced stimulation that reduces charge accumulation and undesired electrochemical reactions at the electrode surface 13. Although the voltage waveforms generated by ME thin films at resonance are biphasic, their resonant frequencies (typically 20-150 kHz) are too high to generate reliable cellular stimulation as mentioned above. To create effective biphasic stimulation within the therapeutic window (100-200 Hz), we use two thin films with different resonant frequencies connected to the same stimulation electrode (panel [a] of FIG. 12). The first thin film is attached to a full-wave rectifier that is oriented to generate positive pulses, and the second thin film is attached to a full-wave rectifier that generates negative pulses.

[0092] The transistors block the current generated by one membrane from propagating through the circuit attached to the other membrane, ensuring that only half of the circuit is active at a time. By switching the magnetic field frequency between the two ME resonant frequencies, we can create a biphasic neurostimulator by alternating positive and negative phase stimulation (Figure 12, panels [b-d]). In this case, a residual charge of -2.3 nC discharging in <2 ms implies that this stimulator can safely operate at frequencies >500 Hz without accumulating charge.

[0093] Using neocortical brain slices from mice expressing the genetically encoded calcium indicator GCaMP3 in GABAergic neurons, we found that the biphasic ME stimulator could repeatedly apply neural stimulation. To image neural activity after ME stimulation, we inserted the stereotrode-mounted biphasic ME stimulator described above while imaging GCaMP activity using fluorescence microscopy (panels [e-g] of Fig. 12, "Methods"). We chose neural stimulation parameters similar to those commonly used for deep brain stimulation

[34] (100 200 Hz biphasic pulses, each phase lasting for 175 μs). When the magnetic field was on, we observed a corresponding increase in fluorescence consistent with activity-mediated calcium increases in n=23 recordings in neocortical layer 5. Following bath application of tetrodotoxin (TTX, 500 nM), the increase in fluorescence was completely blocked in n=9 recordings, confirming that ME stimulation reliably evoked sodium channel-dependent action potentials in nearby neurons.

[0094] In summary, FIG. 12 shows aspects of a biphasic ME stimulator that activates neurons in ex vivo brain slices. Panel (a) of FIG. 12 shows a schematic diagram of the experimental setup using two ME thin films for biphasic stimulation. The setup is similar to the embodiment described in FIG. 1, but includes a microscope objective for observation. Panel (b) of FIG. 12 shows the measured voltage waveform generated by the magnetic field generator. When coupled to a magnetic coil, this waveform generates a magnetic field that alternates between the resonant frequencies of the two ME thin films. Panel (c) of FIG. 12 shows the measured voltage across the stereotrode, which shows the expected biphasic pulse shape. Panel (d) of FIG. 12 shows the calculated current based on measurements of the voltage across the load resistor (VR), showing near perfect charge balance, with only 2.3 nC of charge accumulating on the electrodes per pulse train. Panel (e) of Figure 12 shows a brightfield image (scale bar = 1 mm) of a stereotrode in mouse cortex with an inset of the GCaMP signal averaged over a 600 μm × 600 μm area around the stereotrode tip. The arrows indicate cell bodies fluorescing near the stereotrode. Panel (f) of Figure 12 shows the average GCaMP signal upon application of a resonant magnetic field before (panel [f]) ​​and after (panel [g]) addition of TTX, demonstrating that neural activity is evoked by the ME stimulator. Thin green traces represent separate experiments from two different brain slices, while thick black traces represent the average of all experiments.

[0095] ME neurostimulation in freely moving rats exhibits behavioral efficacy The main advantage of the ME stimulator disclosed herein is the fact that remote activation allows for experiments in freely moving animals. As a proof of principle, we adapted the biphasic stimulator for deep brain stimulation (DBS) in freely moving rats, as shown in FIG. 13. To test the ME stimulator efficacy, we used a previously reported protocol

[36] for testing DBS in hemiparkinsonian rats. In these experiments, rats are injected with 6-OHDA into the left medial forebrain bundle (MFB) to create a unilateral lesion of the substantia nigra pars compacta (SNc). The rats are then placed in a circular enclosure with a diameter of 30 cm. After methamphetamine administration, hemiparkinsonian rats were shown to rotate to the same side as the injection (e.g., left side in case of injection into the left MFB). During these rotations, the rats move mainly using the contralateral (right) forepaw and rarely place the ipsilateral (left) forepaw on the ground. When biphasic stimulation at 200 Hz is applied to the subthalamic nucleus (STN) using a tethered electrode array stimulator, rats typically stop turning to the left and display more normal behavior, such as moving on both forepaws, maintaining a stable orientation, or turning to the other side

[34] .

[0096] To create a wireless biphasic ME stimulator for freely moving animals, we added a small permanent magnet to generate a bias magnetic field to the ME stimulator and wrapped the behavioral chamber with 18 AWG copper wire to create a solenoid (panel [a] of FIG. 13, FIG. 14). By incorporating a small permanent magnet (<0.25 g) into the ME stimulator, we ensured that the bias magnetic field was always aligned with the ME membrane as the rat moved within the enclosure. We also ensured that positive and negative stimuli had equal amplitudes by independently adjusting the distance between each membrane and the permanent magnet. This ME stimulator was then connected to a commercially available electrode array (microprobe) implanted in the STN (panel [b] of FIG. 13, “Methods”). We ensured that the stimulation voltage and current were within a safe therapeutic range by measuring the output of the ME stimulator connected to an equivalent circuit model of the brain (panel [c] of FIG. 13, “Methods”). Specifically, we observe peak voltages of about ±1.5V and peak currents of about ±100μA at a duty cycle of about 50% for 400us (total current per phase of 200μs), which is within the range of effective stimulation reported for conventional wired stimulators

[36] . When we tune the magnetic field frequency off-resonance, we observe almost no generated voltage or current (panel [c] of Figure 13).

[0097] Effective DBS in a freely moving rat using a wireless ME stimulator is shown in Figure 13. Panel (a) of Figure 13 shows the experimental setup showing the rat in a circular enclosure wrapped with magnet wire. The inset shows the biphasic ME stimulator on a penny. Panel (b) of Figure 13 shows a schematic diagram of the biphasic ME stimulator attached to an electrode array implanted in the STN. Panel (c) of Figure 13 shows the measured voltage generated by the ME device and the current applied to the brain in on-resonance (green) and off-resonance (blue) conditions. Biphasic stimulation of approximately 100 μA is applied only when the magnetic field frequency matches the resonant condition. Panel (d) of Figure 13 shows the angular velocity of a hemiparkinsonian rat in a 40-minute DBS trial using intervals of resonant and non-resonant stimulation, showing that rotation is reduced only when the stimulator is activated by the resonant magnetic field. Panel (e) of Figure 13 shows typical trajectories of the rat's head position over the two 30-s intervals shown in c (scale bar = 5 cm). Panel (f) of Figure 13 shows the rat's average angular velocity during the 30 s before and the first 30 s of stimulation for each interval during the 40-min experiment, showing a clear decrease in angular velocity only when the ME thin film was activated in the on-resonance state ( *** P = 4 × 10-7, ns = non-significant P = 0.70, paired t-test).

[0098] We tested a wireless version of the biphasic ME stimulator attached to the head of freely moving rats and found that ME stimulation exhibited efficacy comparable to previously reported wired DBS stimulators. We found that when the magnetic field was applied in a resonant state, a 1-minute period of 200 Hz biphasic pulse resulted in a significant reduction in the rats' rotational speed, as shown in the green interval of panel (d) of Figure 13. This reduction in rotation was not observed when the magnetic stimulation frequency was tuned off-resonance, as shown in the blue interval of panel (d) of Figure 13. Plots of head trajectories show that when the ME stimulator is active during resonant magnetic field stimulation, the pathological rotation observed during off-resonant magnetic field stimulation is absent (Figure 13 panel [e], "Methods"). Averaged across all trials, the mean rotational speed during the first half of stimulation was reduced to a statistically significant 1.6 revolutions per minute (rpm) compared to 9.3 rpm in the absence of stimulation or 9.4 rpm during off-resonance stimulation (paired t-test, FIG. 13, panel [f]). We further demonstrated the reproducibility of this stimulation device by repeating the stimulation protocol on a second rat and found similar results (FIG. 14, panel [b]).

[0099] Panel (a) of Figure 14 shows a schematic diagram showing the position and spacing of the wires and the number of drivers used to generate the alternating magnetic field that overlays a COMSOL simulation of the magnetic field strength inside the chamber. Panel (b) of Figure 14 shows the results of the rotation test in rat 2. The angular velocity for 30 s before stimulation and during the first 30 s of stimulation shows a clear decrease in angular velocity only when the ME thin film is activated in the on-resonance state ( *** P = 1.5 × 10-8, ns = non-significant P = 0.27, paired t-test).

[0100] With a weight of 0.67 g, the ME stimulator described here is the first reported miniature magnetic high frequency stimulator. Moreover, by varying the frequency and timing of the external drive coil, we can generate diverse stimulation patterns across the entire 100-200 Hz therapeutic window, which is also applicable to other disease models. In addition, calculations of the magnetic field strength suggest that we can reconfigure the drive coil for numerous behavioral experiments by placing the coil under the floor of the animal enclosure. Finite element simulations and measurements show that even at a distance of 4-5 cm above the drive coil, the ME thin film generates sufficient voltage for stimulation (Figure 15). This distance can be further improved by optimizing the coil geometry or increasing the power of the magnetic field. Figure 15 shows that the ME device can operate several centimeters above a single magnetic coil. Panel (a) of Figure 15 shows a COMSOL simulation of the magnetic field above a circular coil, and panel (b) of Figure 15 shows the measured device output voltage as a function of distance above the coil.

[0101] Demonstration of multi-channel deep brain stimulation in a skull model using rice-grain-sized ME thin film In addition to supporting experiments in freely moving rodents, ME materials can enable miniaturized wireless stimulators operating deep in the brain of larger animals or human patients and individually activated by external electromagnets. To our knowledge, this is the first technology to enable independent external wireless control of multiple miniature stimulators deep inside a human skull model. Figure 16 shows the predicted depth at which various miniature antennas can be safely implanted under the skull and generate 1 mW of power, the approximate maximum power required for high frequency continuous neural stimulation

[31] . As mentioned above, radio frequency (RF) powered antennas operating at frequencies above about 1 MHz are limited in the amount of power that can be safely applied to the implanted device without causing potentially harmful tissue heating. Simulations show that when operating at safe power limits, in order to harvest 1 mW of power, the RF antenna must be placed on the surface or in very shallow regions of the brain. "Mid-field" techniques

[37] improve RF coupling efficiency to enable deep operation, but because this technique operates at a fixed frequency, individually addressable motes or biphasic stimulation have not yet been demonstrated. Other techniques for wireless power delivery mentioned above, such as magnetic induction, also cannot achieve deep multichannel stimulation. For example, even using a higher operating frequency of 1 MHz, an induction coil with the same orientation and cross-sectional area as the ME thin film shown here would require a minimum of 500 turns of wire (assuming a typical Q factor of 10) to generate 2 V using the same 0.5 mT magnetic field used here. Thus, devices based on magnetic inductors cannot be miniaturized without sacrificing available power as mentioned above [9].

[0102] Figure 16 shows aspects of miniaturized multi-channel stimulation in a human skull model. Panel (a) of Figure 16 shows the effective depth under the human skull model for the ME device compared to other miniature wireless stimulators. The depth limit is based on the safety limit for generating 1mW (i: Park et. al, Proc Nat Ac Sci, 2016; ii: Yazdandoost et. al, Asia Pac Microw Conf, 2009; iii: Yazdandoost et. al, Proc 37 Europ Microw Conf, 2007; iv: Agrawal et. al, Nat Biomed Eng, 2017). Panel (b) of Figure 16 shows a photograph of the ME thin film next to a grain of rice and the corresponding voltage as a function of magnetic field frequency (field strength 1mT, scale bar 2mm). Panel (c) of FIG. 16 is a schematic diagram showing a potential application of a fully embedded ME thin film using a magnetic field generated by an external coil that can be integrated into a hat or visor. Panel (d) of FIG. 16 shows a front view of the skull model with the top removed to reveal the LEDs, and panel (e) shows its top view (arrows indicate the location of the thin film; scale bar 1 cm). Panel (f) of FIG. 16 shows a photograph of the LEDs attached to the ME thin film subjected to a magnetic field at 180 kHz, and panel (g) of FIG. 16 shows a photograph of the LEDs attached to the ME thin film subjected to a magnetic field at 200 kHz. Selective illumination of the LEDs corresponding to the resonant frequency of the thin film demonstrates successful multi-channel activation of individual thin films (scale bar 1 cm). The magnetic field strength was measured to be 0.5 mT at the location of the ME thin film.

[0103] As a proof-of-concept experiment, we show that two rice-grain-sized ME membranes can be individually addressed at the center of a human skull model using an external electromagnet. These two membranes, 8 mm and 10 mm long, have acoustic resonance frequencies of 180 and 200 kHz, determined by the length of the membrane. When these membranes are attached to orange LEDs, their output voltage is capped at about 1.8 V, which helps control the stimulation voltage and allows membrane activation to be visualized. ME membranes of this size are smaller than current DBS leads and can potentially be implanted in deep brain regions, as shown in panel (c) of Figure 16. In addition, the magnetic stimulation coil is small enough to be incorporated into a stylish hat or visor that can be comfortably worn by the patient. We found that when two ME membranes are placed at the center of a skull model, applying a magnetic field at the resonant frequency of the selected membrane can individually light up an LED on each membrane (Figure 16 panels [d-g]). For this experiment, we used a 400W power supply, which generated a magnetic field of about 0.5mT at the center of the skull model. The top of the skull model was removed for visualization, but this did not affect the ability to drive the LED indicators. The number of stimulation channels can also be increased by adding ME films with different resonant frequencies.

[0104] To our knowledge, this is the first demonstration of a miniature magnetic nerve stimulator that (1) operates at therapeutic bands (100-200 Hz) in freely moving animals; and (2) allows for individually addressable miniature stimulators deep within a human skull model. However, the advantages of ME materials go beyond these proof-of-principle demonstrations.

[0105] ME stimulators, such as the examples described in the in vivo rat experiments, can have a direct impact on DBS therapy research using rodent disease models. Because the ME stimulator is compatible with commercially available implanted electrodes and the magnetic stimulator can be adapted to many standard behavioral experiments or animal enclosures, the ME stimulators disclosed herein can easily replace currently used wired DBS stimulators. As a result, new experiments can be developed to probe the effects of DBS in social situations where long-term, continuous DBS or wired DBS stimulators are not feasible.

[0106] In addition, ME materials have the potential to enable miniaturized neurostimulators that can be implanted deep in the brains of larger animals or humans and addressed externally with small electromagnets. As shown here, thin films the size of a grain of rice can be selectively activated based on their unique resonant frequency. Because the voltage depends on the thickness of the piezoelectric field and not on the length of the film, further miniaturization is not expected to reduce the voltage generated by these films (panel [c] in Figure 7), suggesting that even smaller films can act as effective stimulators.

[0107] We also foresee the application of ME materials as a wireless powering technology for more complex implantable bioelectronic devices. For example, the demonstrated ability of ME thin films to power LEDs implies that ME materials could also power implantable optogenetic stimulators or tiny integrated circuits for physiological monitoring.

[0108] To realize these fully implantable bioelectronic devices, work is needed to improve ME materials and manufacturing processes to reliably produce high-quality miniature ME thin films and encapsulate them for long-term use. For wearable technology, magnetic field generators need to be further miniaturized so that they can be battery-powered and comfortable to wear. These advances must also be accompanied by in vivo testing to demonstrate safety and efficiency for long-term use.

[0109] Collectively, ME materials have the potential to meet the critical need for wireless powering of miniature biphasic neurostimulators and other bioelectronic devices, where the main challenge is to transmit energy over distances of several centimeters without heating the tissue or suffering losses at the interfaces between tissue, bone, and air.

[0110] method General statistics: Error bars in panels (f) and (g) of Figure 7 indicate ±1 standard deviation of n=-50 data points. We further performed Tukey's HSD test on the data in panel (g) of Figure 7, which showed that the voltages generated for different PVDF thicknesses were significantly different. A paired t-test was used for the rotational test in panel (e) of Figure 13.

[0111] thin film manufacturing To fabricate the ME thin film, we used Metglas SA1 alloy (Metglas Inc) for the magnetostrictive layer and polyvinylidene fluoride "PVDF" (precision acoustics) for the piezoelectric layer. The PVDF thin film used in these experiments was pre-stretched and poled by the manufacturer. The two layers were bonded together using an epoxy (Hardman double bubble red epoxy) that can transmit mechanical stress between the two layers. Before bonding the two layers, we sputtered a thin platinum layer (<100 nm) onto the PVDF as the top electrode. Before the epoxy treatment, O 2Both Metglas and PVDF were plasma cleaned for 5 minutes using plasma. After epoxy curing, scissors were used to cut the film into the desired rectangle, taking care to cut the long axis of the film along the stretch direction of the PVDF. We then attached wires to each side of the film using conductive epoxy to measure the electrical capabilities of the film. We found that attaching the wires to the center dramatically increased the resonant voltage. However, for convenience, the wires were attached near the ends of the film during in vitro experiments. In many cases, we also attached additional electronic components, e.g., diodes or LEDs, to the wires attached to the film, as noted in the appropriate sections of the text. Finally, the devices were coated with 5-10 μm of Parylene C (Labcoater 2). Initially, this coating was used to electrically insulate and protect the device during in vitro experiments, but we found that encapsulation increased the resonant voltage, which may have been due to the increased mechanical bonding provided by the encapsulation.

[0112] Bench-top electrolysis test To observe bubble formation from electrolysis at the tip, the stimulator shown in panel (a) of FIG. 13 was hardwired to a stereotrode immersed in saline under a microscope. As demonstrated above, during monophasic stimulation we used only one resonant frequency, and during biphasic stimulation we used two frequencies. In both cases the pulse duration was 400 μs / phase. We determined the limit of stimulation time as when the first bubbles began to appear at the tip of the electrode, and each data point was repeated four times.

[0113] Magnetic field generation In the embodiment shown in FIG. 9, each magnetic field generator contains two main components: (1) a magnetic coil used for the alternating magnetic field (described in the text and optimized for each experiment); and (2) an electronic driver (same for all experiments) to control the voltage and timing of the alternating current in the coil.

[0114] To maintain simplicity, efficiency and low cost, the coil was driven with a full H-bridge style switching circuit. The driver is designed to deliver high current to the drive coil in the form of a biphasic pulse train. This reduces the cost and complexity of the driver itself as well as the power supply and control circuitry compared to an arbitrary function generator. This design also has the ability to improve operational efficiency through impedance matching with the drive coil. In addition, it is also possible to control the power delivered to the drive coil on the fly by adjusting the duty cycle of the current pulses, allowing for easy numerical control of the power delivered to the ME thin film while maintaining a resonant carrier frequency. The magnetic field output carrier and pulse frequency are generated using a TeensyLC board and custom Arduino code to generate specific pulse timing to deliver controlled ME stimulation (panels [c] and [d] in Figure 9).

[0115] These coils and drivers can be combined in various ways to generate magnetic fields appropriate for a given experiment. For example, the apparatus used to generate alternating magnetic fields in the in vivo rotation experiments included four sets of coils, each with five turns, powered by one driver when all four drivers were synchronized to the same output signal. In this way, we can generate enough power to generate mT-scale magnetic fields throughout the entire activity zone (panel (a) of FIG. 14).

[0116] cell culture For experiments performed on coverslips, the sodium channel Na 1.3 and potassium channel K 2.1HEK cells expressing ArcLight were grown on 12 mm poly-l-lysine-coated coverslips to approximately 30% confluency. Cells were then transfected with the genetically encoded voltage indicator ArcLight using Lipofectamine (Invitrogen) according to the manufacturer's recommendations. Two to three days after transfection, coverslips were inverted onto ME thin films in preparation for testing. The preparation of GFP controls followed the same procedure, except that a GFP expression vector (AddGene) was used instead of the ArcLight vector (AddGene). For experiments performed with cells grown on thin films, ArcLight-transfected HEK cells were placed on parylene-coated poly-l-lysine-treated thin films. Thin films were placed in cell culture medium overnight and tested the next day.

[0117] ArcLight and GFP were excited at 460 nm by an LED light source. Fluorescence images were collected at 33 fps using a CCD camera. Images were analyzed using Matlab to quantify fluorescence changes in individual cells. Extracellular buffer (ECB, in mM: NaCl 119, KCl 5, Hepes 10, CaCl 2 2,671 MgCl 2 In vitro studies were performed in 10 mL of 100% ethanol (pH 7.2; 320 mOsm).

[0118] Panel (b) of Figure 11 was obtained by growing unmodified HEK cells on the membrane immersed in cell culture medium for 5 days. The cells were then stained with Hoechst and Calcein-AM to label the nuclei and membranes, respectively, in live cells. The cells were then fixed and imaged using a confocal microscope.

[0119] Mouse brain sectioning procedure We used 40-day-old GAD2-GCaMP3 mice, which were generated by crossing GAD2-Cre (JAX#10802) mice with floxed-GCaMP3 (JAX#14538) mice. Preparation of brain slices followed the procedure described by Ting et al.

[45] and was performed in accordance with the National Institutes of Health guidelines and approved by the THealth Animal Welfare Committee. Mice were deeply anesthetized with isoflurane and maintained in a 10-mL bronchodilator containing 92 mM NMDG, 2.5 mM KCl, 1.25 mM NaHCO3, and 1.0 mM ethanol. 2 PO 4 , 10 mM MgSO 4 , 0.5 mM CaCl 2 , 30 mM NaHCO 3 , 20 mM glucose, 20 mM HEPES, 2 mM thiourea, 5 mM sodium ascorbate, 3 mM sodium pyruvate, 95% O 2 and 5% CO 2 Coronal brain sections (300 μm) were cut using a vibratome (Leica VT1200S) and incubated in NMDG solution for 15 min at 35°C, followed by perfusion with 92 mM NMDG, 2.5 mM KCl, 1.25 mM NaHCO, and 1.0 mM NaCl. 2 PO 4 , 2 mM MgSO 4 , 2 mM CaCl 2 , 30 mM NaHCO 3 , 25 mM glucose, 20 mM HEPES, 2 mM thiourea, 5 mM Na ascorbate, 3 mM Na pyruvate, 95% O 2 and 5% CO 2 For the experiments, slices were transferred to a chamber saturated with 126 mM NaCl, 2.5 mM KCl, 1.25 mM NaHCO, and maintained at room temperature. 2 PO 4 , 2 mM MgCl 2 , 2 mM CaCl 2 , 26 mM NaHCO 3The stereotrode was placed in a recording chamber perfused with ACSF containing 10 mM glucose and maintained at 32-34 °C using a 695 in-line heater. NBQX (10 µM) was included in the bath solution to block AMPA receptor-mediated synaptic transmission. The stereotrode was placed in layer 5 of the somatosensory (barrel) cortex.

[0120] GCaMP3 was excited at 460 nm by an LED light source. Fluorescence images were collected at 9.8 fps using a CCD camera attached to an Olympus BX51WI microscope. Images were analyzed using Matlab to quantify the fluorescence changes in a 600 × 600 μm area around the stereotrode chip.

[0121] Implant design and rat surgical procedures Two male Long-Evans rats (n ≈ 1, 400 g) were anesthetized with isoflurane gas. Anesthesia was induced using 5% isoflurane and 2% was used to maintain anesthesia depth. Buprenorphine (0.04 mg / kg) for analgesia was administered 30 min before ear bar. Five to seven skull screws were placed to fixate the electrode array. The skull screws were bonded to the skull by Metabond dental acrylic resin. A craniotomy was performed to accommodate the microelectrode array and expose the injection site for the neurotoxin. The dural covering of the brain was incised and peeled away with a 30-gauge needle with a bent tip. To protect noradrenergic neurons, desipramine (DMI) reconstituted at a concentration of 15 mg / mL in saline was injected intraperitoneally. The dose of DMI was approximately 15 mg / kg and was injected approximately 30 min before neurotoxin administration. To induce hemiparkinsonian lesions, 8ug of 6-hydroxydopamine (OHDA) at 2μg / μL in saline was injected at 0.2uL / min into the mid forebrain bundle (MFB, -1.2ML, -4AP and -8.1DV). STN stimulation was delivered via a 2×2 platinum-iridium microelectrode array (microprobe) with 75μm electrodes spaced 600×600μm apart. Each electrode had a nominal impedance of 10kOhm. The electrode array was lowered from bregma to -2.6ML, -3.6AP and approximately -8.2DV. The array was fixed to the skull with dental acrylic. All experiments were approved by the Institutional Animal Care and Use Committee of Rice University.

[0122] Prior to stimulating each rat with the magneto-electrical stimulator, stimulator power was estimated by a benchtop approximation of rodent electrode impedance. Constant current stimulation of the rodent brain with an AM Systems 4100 stimulator produced a characteristic voltage waveform that approximated a simplified parallel RC circuit. A 56 kOhm resistor and a 440 pF capacitor in parallel closely approximated the impedance characteristics of the rat brain across the stimulating electrodes. Using this circuit model, we estimated the field strength and pulse duration required to produce the desired stimulation effect and ensured that the stimulation was charge-balanced prior to rodent experiments.

[0123] Rotation Test Before performing the rotation test, rats were briefly anesthetized with 5% isoflurane gas, intraperitoneally injected with methamphetamine (0.31 ml, 1.25 mg / kg), and a wireless biphasic stimulator was connected to the implanted electrode array. After the anesthesia wore off (approximately 5–10 min), the rat was placed in a cylindrical behavioral chamber. A magnetic field was applied across the entire behavioral area to the thin film on the device (panel [a] in Figure 14).

[0124] The magnetic field was applied on and off resonance for 1 min at various times during the 40 min trial. The resonance frequencies were 73 kHz and 747 77 kHz, and the off resonance frequencies were 63 kHz and 87 kHz.

[0125] Rodent Tracking To track the ears, proboscis and implants, we used a slightly modified version of DeepLabCut

[46] to generate head positions on the rotation task. We hand-labeled a dataset of 286 frames in total from both the on-resonance and off-resonance rotation tasks and trained for approximately 140,000 iterations.

[0126] Skull model demonstration experiment Since bone and tissue are essentially transparent at the magnetic field frequencies used in this experiment

[47] , we selected a life-size phantom with an average adult head size as the skull (Orient Infinity Limited). It was wrapped with 18 AWG magnet wire as shown in Figure 16. The coils included four coils in parallel, each hardwired to an individual magnetic field driver. All drivers were hardwired to the same input frequency signal and powered from the same power source. The membrane was suspended in the center of the skull phantom. An orange LED (Chanzon) with a diode in anti-parallel was attached to the membrane for wireless verification of the voltage generated by the membrane. The top of the skull was removed to better photograph the LED for visualization purposes.

[0127] Zinc oxide (ZnO) As previously mentioned, certain embodiments may include a thin film doped with zinc oxide (ZnO) nanoparticles. Referring now to Figure 17, a graph shows the voltage over time for a PVDF / Metglas thin film (about 8 μm) doped with ZnO nanoparticles, resonating at about 70 kHz. Upon leaving the resonant frequency, the voltage drops to zero.

[0128] Manufacturing method In preliminary trials, PVDF-TrFE (a modified PVDF molecule with relatively low temperature sensitivity) was dissolved in the solvents methyl ethyl ketone and butanone at 20 wt%, following a similar procedure from Dodds et al.

[49] . ZnO nanoparticles (diameter about 13 nm) were added to the solution at about 10 wt%, and the solution was placed in a bath sonicator for more than 6 hours. The solution was then spin-coated twice at about 3000 RPM for 45 seconds onto a 3 x 3 cm piece of Metglas to obtain a film thickness of about 8 μm. The laminate was then thermally annealed at 150 °C and electrically poled in a custom-made corona poling apparatus. This resulted in one batch of four thin films with higher voltage and self-rectifying properties (see Figure 17). Challenges included reproducibility of results, where ZnO nanoparticles are insoluble in the PVDF solution, especially at higher wt%. To the best of our knowledge, there is no evidence in the literature showing the self-rectifying behavior of PVDF-TrFE mixed with nanoparticles.

[0129] PVDF-ZnO nanowire materials are known that reliably exhibit self-rectifying behavior.

[50] Since the ZnO nanowires are grown in situ, they act as self-regulating diodes to control the output of the hybrid material.

[0130] In a particular preparation, the concentration of PVDF in the solvent remains the same (20 wt% in MEK and butanone). Instead of using ZnO nanoparticles, a particular embodiment may grow ZnO nanowires from ZnO nanoseeds. A possible embodiment may use Zn(Ac2) (zinc acetate) as a seed and mix it with the PVDF solution. The hybrid solution can then be spin-coated onto a Metglas substrate.

[0131] In a particular embodiment, the thin film can be placed in a ZnNO3 (zinc nitrate) aqueous solution and the nanowires are allowed to grow naturally at 80°C for 6 hours. The timing and end of the reaction depends on the size of the nanowires desired. Furthermore, since PVDF is naturally stressed by the extensional forces due to ZnO NW growth, the PVDF material naturally forms into the piezoelectric beta phase. However, if the poling process is insufficient, we can also electrically pole the material with a corona poling device. The final product is probed and tested with a microwire to measure the voltage at the resonant frequency. SEM imaging can also be used to verify the ZnO NW growth in the thin film and X-ray diffraction can be used to verify that the thin film is in a beta piezoelectric conformation. Figure 18 shows X-ray diffraction data showing the poling of PVDF with the α and β phase peaks of PVDF, which can be used to verify the thin film after ZnO nanowire growth.

[0132] Figure 19 shows aspects of ZnO / PVDF thin film fabrication. Panel (a) of Figure 19 shows that zinc acetate is incorporated into the PVDF solution to act as nanoseeds, which form a thin film in zinc nitrate to grow nanowires. Panel (b) shows a scanning electron microscope (SEM) image of ZnO nanowires grown in PVDF. Panel (c) shows an SEM image of ZnO nanowires on the surface of PVDF, and panel (d) of Figure 19 shows a close-up of the ZnO nanowires.

[0133] FIG. 20 shows the self-rectification of electrical current in a PVDF / ZnO nanowire hybrid material used as an energy harvester when the material is pressed with a finger.

[0134] 21 , an apparatus 200 includes an implantable wireless neurostimulator 210 and a magnetic field generator 220 configured to generate a magnetic field 236 at one or more input frequencies. In the illustrated embodiment, the implantable wireless neurostimulator 215 includes a magnetoelectric film 230, a charge storage element 235, an electrical circuit 240, a first electrode 251 and a second electrode 252. In this embodiment, the electrical circuit 240 is coupled to the magnetoelectric film 230 via the first and second electrodes 231 and 232. An exemplary embodiment of the present disclosure may also be referred to herein as a "Mag-Dust" or "MagMote" device.

[0135] In certain embodiments, the magnetoelectric film 230 is configured to generate an electrical signal 245 when stimulated by a magnetic field of a particular input frequency. The electrical circuit 240 can be configured to modify the electrical signal 245 to charge the charge storage element 235, which in certain embodiments can be a capacitor. Additionally, the electrical circuit 240 can be configured to transmit data to components of the implantable wireless neurostimulator 210 and can be configured to provide a stimulation output 258 to the third electrode 251 and the fourth electrode 252. In the illustrated embodiment, the stimulation output 258 can be used to stimulate neural tissue 290, and in certain embodiments, the third electrode 251 and the fourth electrode 252 can be configured as a stereotrode.

[0136] 22A and 22B, an exemplary embodiment is shown including certain components of the electrical circuit 240. It will be understood that this embodiment is merely an example and other embodiments of the electrical circuit 240 may include different components or different arrangements of components than those shown in FIGS. 22A and 22B. In this embodiment, the electrical circuit 240 includes a charging module 241 configured to provide power management and charging of the charge storage element 235. In addition, the embodiment of the electrical circuit 240 shown in FIG. 22A includes a data module 242 configured to provide data transmission and control. The embodiment of the electrical circuit 240 shown in FIG. 22A also includes a stimulus driver module 243 configured to provide a stimulus output 258 to the third electrode 251 and the fourth electrode 252. In this embodiment, the charging module 241 includes a rectifier, a reference generator, a low dropout (LDO) linear voltage regulator, and a power on reset (POR). The illustrated stimulus driver module includes a reference current generator and a current digital-to-analog (DAC) converter. In addition, the illustrated data modules include a control signal generator, a stimulus controller, and a clock recovery module. Figure 22B shows a circuit diagram of a specific embodiment of a rectifier, a native MOS-based voltage reference, and a 5-bit programmable stimulus driver. Further description of the operational aspects of the electrical circuit 240 and specific components are provided below in the section entitled "Examples."

[0137] In an exemplary embodiment, the electrical circuit 240 includes a complementary metal oxide semiconductor (CMOS) chip 247 that can receive power and commands (e.g., through data transmission) and interface with the ME transducer 239 incorporating the magnetoelectric thin film 230 to drive the programmable electrical stimulation. In the illustrated embodiment, the AC voltage induced on the magnetoelectric thin film 230 is first converted to a DC voltage by an efficient active rectifier driven by a four-input differential comparator. Energy can be temporarily stored in the charge storage element 235 in preparation for the next stimulation phase. In a particular embodiment, a low dropout regulator provides a stable 1V supply to the rest of the system (except for the stimulation) to reduce their power consumption and effectively increase the load impedance of the application specific integrated circuit (ASIC).

[0138] As shown in FIG. 23, the ASIC can cycle through a charging step 261, a data receiving step 262, and a stimulating step 263. In a particular embodiment, downlink data transmission is performed by amplitude shift keying (ASK) of the magnetic field (shown in FIGS. 21 and 22A). The received data can be used to program the system operation and stimulation modes. The built-in driver stimulation can be used to drive the electrodes 251 and 252 (shown in FIGS. 21-22) with current-regulated biphasic pulses that cover a wide range of stimulation patterns appropriate for therapy (e.g., frequencies from 0-200 Hz, amplitudes from 30-1500 μA, and pulse widths from 50-500 μs).

[0139] One challenge to the ASIC design is to maintain reliable operation even as the amplitude of the voltage induced on the ME transducer 239 changes under different distances and angles between the ME transducer 239 and the magnetic field generator 220. Exemplary embodiments can address this by maximizing the power conversion efficiency of the device, making the system adaptable to environmental conditions, controllable by an external transmitter, and calibration-free. These properties can be achieved through three circuit and system features: (1) an optimized power interface to the ME transducer using different materials; (2) an adaptive ME-based data receiver to reliably initiate data transmission and decode data even when the ME-induced voltage changes; and (3) low-power and calibration-free circuitry to minimize the quiescent power of the ASIC and maximize the energy available for stimulation.

[0140] CMOS Power Interface to ME Transducers Increasing the power conversion efficiency of Mag-Dust to extract more usable power from low-strength magnetic fields is crucial to realizing our goals of (1) reliable operation under large physical distance and angle changes, and (2) a wearable battery-powered transmitter. To miniaturize the implant, Mag-Dust does not incorporate a battery and is continuously powered by the external magnetic field. Within each cycle, all harvested energy is either used for stimulation and system operation or wasted. As a result, it is more efficient to extract high voltage directly from the ME transducer than to pursue the maximum power point by a low-efficiency DC-DC up-converter as in most ambient energy harvesters. To maximize the usable power in Mag-Dust, we study the electrical model of the ME transducer, the power management of Mag-Dust, and the impedance matching between the ME transducer and the ASIC.

[0141] The electrical model of the transducer is useful for interface design and transducer-circuit co-optimization. The magnetoelectric laminate composite can be modeled with an RLC model shown in panel (a) of Figure 24. This model can be calibrated with COMSOL physical simulations and measurements of nanomachined samples. As shown in panels (b) and (c) of Figure 23, the ME transducer nanomachined with PZT piezoelectric material induces peak voltages at the mechanical resonance of 240 kHz and two purely resistive effective electrical resonance frequencies. It should be noted that some ME materials (e.g., PZT) exhibit a purely resistive impedance at resonance, similar to ultrasonic transducers

[10] , while other ME materials (e.g., PVDF) are capacitive in the vicinity of resonance.

[0142] The development of the exemplary embodiment has led to significant observations regarding wirelessly powered battery-less devices. In particular, the maximum load power (e.g., in most cases, the stimulation phase) is equal to the effective recharge power for charging the storage element 235 (as shown in Figs. 21-22). In contrast to the traditional power transfer model, where increasing load impedance reduces the received power, the available power stored in the charge storage element actually increases with the received voltage. Preliminary simulation results (based on the ME model of Fig. 24) confirm this behavior and imply the importance of reducing the effective load impedance of the ASIC by reducing the quiescent power, as shown in panel (a) of Fig. 25. Fig. 25 shows the available power under impedance matching between the ASIC and the ME transducer, where panel (a) shows a purely resistive impedance and panel (b) shows a capacitive impedance. Compared to attaching the ME transducer directly to the electrode with an impedance of several kilo-ohms, the effective load impedance of the ASIC can be as large as tens of kilo-ohms in certain embodiments.

[0143] Consistent with previous work on PVDF materials, our preliminary measurements of a mm-sized ME thin film based on PVDF show a capacitance of 68 pF at 150 kHz resonance. Because lead-free PVDF is potentially more biocompatible than PVT, we study circuit interfaces to PVDF-based ME transducers to accommodate their different properties from other commonly found energy harvesters. One challenge is to compensate the capacitive impedance of the ME transducer to maximize power harvesting. One way to achieve this is to add a series inductor. However, as shown in panel (b) of Figure 25, this is a challenge because a millihenry inductor is required to achieve the goal, and inductors of such value are only possible with bulky discrete components. To solve this challenge, we explore the principle of emulating an inductor with a multiphase flipping capacitor for impedance compensation

[41] . We investigate its circuit implementation for PVDF-based embodiments, as well as the reduction of the required on-chip capacitance by adding more phases and exploring new switched-capacitor topologies.

[0144] Adaptive Data Reception When high data rates are not required, amplitude shift keying (ASK) is suitable for low-power data communication to implanted devices. Because the large inductors required for mT strength magnetic field generators prevent fast switching between different amplitudes, certain embodiments may indirectly modulate the voltage induced in the ME transducer by modulating the frequency of the AC magnetic field. However, employing this amplitude modulation scheme presents three challenges: (1) large energy storage capacitors that limit the switching speed of the rectified DC voltage; (2) variations in the resonant frequency and voltage conversion efficiency of the ME transducer due to manufacturing variations; and (3) unpredictable ME-induced AC amplitudes that depend on physical conditions, such as the distance and alignment between the ME transducer and the transmitter.

[0145] To avoid large off-chip capacitors during data reception, the exemplary embodiment may use a dual-path (one for power supply and one for data reception) structure. We propose to control the switching between the two paths by a watchdog logic that monitors the change in ME induced voltage. As shown in panel (b) of FIG. 23, the transmitter is designed to reduce the average amplitude of the ME induced voltage during data transmission. In fact, the exemplary embodiment leverages the existing comparator in the active rectifier as a zero-overhead mode switching watchdog. If the amplitude of the input is lower than the voltage stored in the energy capacitor, the comparator can generate a train of narrow pulses. The exemplary embodiment can extract the envelope of this pulse train as an enable signal for data transmission, as shown in panel (b) of FIG. 26. In this way, the external transmitter (e.g., the magnetic field generator 220 shown in FIGS. 21 and 22) has control over the duration of the charging and data transmission steps.

[0146] To synchronize data sampling clocks and thresholds between transmitters and implants under manufacturing and environmental variations, state-of-the-art miniature implants all rely on individually characterizing each manufactured device and calibrating the transmitter to it [10-12]. This approach suffers from three major drawbacks: (1) testing and storing these settings increases the implementation complexity and requires pairing of the transmitter with the implant; (2) precise matching is not feasible because the on-chip frequency depends on the supply voltage and aging conditions, leading to potential data errors and erroneous commands; and (3) the ASKed signal can shift or become distorted under various physical conditions, and a fixed threshold voltage cannot reliably distinguish between "1" and "0". To address these limitations, exemplary embodiments can include a simple and effective method for extracting both the clock and threshold voltage from an external AC magnetic field. The clock can be directly extracted from the frequency of the AC magnetic field, so the frequency of the recovered clock is exactly the same as the frequency of the transmitter. Exemplary embodiments can also derive the threshold voltage from a toggling pilot tone sent by the transmitter before each data transmission session. As a result, the threshold voltage can be set to 50% of the actual receive amplitude at run time, which is otherwise unknown to the system. In a circuit implementation, the received pilot signal can first be low-pass filtered to extract its DC component, and then sampled by a track-and-hold circuit as the final threshold voltage for data decoding, as shown in panel (b) of FIG. 26. A simulated operation of the adaptive data reception scheme is shown in FIG. 27.

[0147] Low power and calibration-free circuit components Exemplary embodiments of the present disclosure may also include a low-power, calibration-free voltage reference for the power management circuitry, a current reference for the stimulus driver, and precise stimulus timing control. As previously discussed, reducing the quiescent power of the exemplary embodiment ASIC is important to increase the available power under a particular magnetic field. In addition, calibration-free circuits are desirable to reduce implementation complexity and improve robustness to environmental variations. Although the results from this exercise are primarily designed for the exemplary embodiment ASIC, the techniques may be applicable to other low-power applications.

[0148] Calibration-free voltage and current references: Conventional voltage and current references using the band gap of bipolar transistors have high bias power, hindering their application in battery-less embedded devices. Threshold voltage based references have emerged for these applications due to their simplicity and low power

[42] . However, a major drawback of existing designs is that they require multi-point calibration to reach good supply and temperature stability. In an attempt to eliminate calibration, a key observation made in PI Yang's previous work is that body bias can be used as an additional tuning knob to compensate for temperature effects [27,43]. Voltage and current references designed with this principle show a significant reduction in manufacturing variation over the prior art, as shown in panels (a-c) of Figure 28. However, their supply voltage sensitivity is still not good enough. Voltage stability is more important than temperature stability because the supply voltage in the exemplary embodiment is sensitive to distance and angle changes. The exemplary embodiment of the present disclosure leverages conventional native transistor stack technology

[18] to obtain significantly improved voltage stability at the expense of some temperature stability. Figure 28 shows low power CMOS circuit components. Panels (a-b) show a calibration-free voltage / current reference, panel (c) shows measurements including manufacturing variations, panel (d) shows a programmable stimulus driver, and panel (e) shows a Wheatstone bridge-based timing reference.

[0149] To achieve robust programmable stimulation modes, the exemplary embodiment may utilize a finite state logic machine with a clock recovered from the AC magnetic field to control the digital-to-analog current converter and the H-bridge output driver stage, as shown in FIG. 28 panel (d). It is expected that the width and frequency of the stimulation pulses can be numerically adjusted to the desired values ​​even as the resonant frequency of the ME transducer varies. However, if the precision of the stimulation timing control proves insufficient, the exemplary embodiment may utilize a CMOS RC oscillator, rather than a bulky quartz oscillator, to provide a timing reference without increasing the volume and weight of the implant. The use of a Wheatstone bridge, as shown in FIG. 28 panel (e), can eliminate the high bias power presented in state-of-the-art designs. A prototype of the exemplary embodiment shows that this design can be two times more efficient than the state-of-the-art

[44] . We further explore switching of both resistors and capacitors in the bridge to support the programmable and low-frequency therapy-appropriate stimulation modes desired in the exemplary embodiment.

[0150] Millimeter-sized stimulator and in vitro testing As disclosed herein, an exemplary embodiment of a neurostimulator demonstrates wireless power and data transmission via AC magnetic field, efficient interfacing between ME transducer and CMOS ASIC, and totally code-free stimulation control. An exemplary embodiment of the system can include a nanofabricated ME transducer, a CMOS ASIC, a ceramic capacitor for temporary energy storage, and a pair of electrodes (on a flexible PCB). Panel (a) of FIG. 30 shows an exploded view of the exemplary embodiment, and panel (b) of FIG. 30 shows an assembled view of the same embodiment. The components are integrated and fully encapsulated on a polyimide-based flexible printed circuit board. Challenges in demonstrating a proof-of-principle system include its functionality as well as its ability to maintain biocompatibility and safety in a biological environment. We will investigate: (1) an encapsulation and thermal management strategy for the exemplary embodiment, (2) a portable and wearable magnetic field generator for transmitting power and data to the exemplary embodiment, and (3) accelerated aging and in vitro testing to verify functionality, safety, and efficiency.

[0151] Biocompatible integration In embodiments for biological applications, the transducers and ASICs can be encapsulated in a way that preserves their mechanical resonance and electrical properties in ionic solutions (such as those in the body). Parylene C is commonly used as a passivation layer for implanted electronics due to its chemical inertness and ability to act as a strong ionic barrier. However, damage to the parylene layer is common due to moisture diffusion. To solve this problem, exemplary embodiments may also be fabricated using atomic layer deposition to deposit aluminum oxide, an excellent moisture barrier. In certain embodiments, bilayer thin films use chemical vapor deposition and atomic layer deposition, both of which are conformal processes that allow the ME thin film to be uniformly coated. The combination of aluminum oxide and parylene has been shown to significantly improve the lifetime of implanted neural interfaces in accelerated aging tests

[45] . Although this strategy is an original approach, the use of Al 2 O 3The combination of Parylene C / Al may be insufficient to maintain the mechanical and electrical properties of the magnetoelectric material. 2 O 3 In addition to testing laminates, we also systematically investigate alternative hard and soft materials, such as silicon carbide (SiC) and polyisobutylene (PIB) (see materials overview in Figure 30 for reference).

[0152] In an exemplary embodiment, the thin film is first encapsulated with a soft material to maintain mechanical resonance, and then wrapped in a hard material. Some of our preliminary work has shown that a thin film encapsulated in a semi-soft material, polyimide, retains more than 95% of its maximum voltage. In a particular embodiment, the external components, integrated circuits and capacitors are potted with medical epoxy, a material commonly used for such applications [46-52]. First, the exemplary embodiment uses a wired PCB in solution testing to test the chip and system functionality to probe for test points, and then tests a wireless version of the device.

[0153] thermal management Testing of the exemplary embodiment shows results that are well within the range of power levels that do not cause significant heating of biological tissue. Based on preliminary device impedance and voltage, the total power delivered to the thin film is expected to be 1.5 mW. This power dissipation rate is comparable to DBS stimulators that do not cause significant heating of the brain [45, 46]. Nevertheless, the temperature change caused by the device during long-term in vitro stability testing is directly measured. Using a fiber optic thermometer, the temperature of the solution can be monitored to ensure that it does not rise more than 2°C, which is considered a safe level for neural implants

[47] . If the temperature is found to rise more than 2 degrees Celsius during testing, lower power stimulation paradigms will be investigated by reducing the duty cycle or improving device characteristics (e.g., reducing electrode impedance) to improve power efficiency.

[0154] Wearable Magnetic Transmitter An advantage of the exemplary ME thin film as a wireless stimulator is that the magnetic fields required to activate these devices require very little power compared to other similar wireless power delivery technologies. For example, the inventors have been able to successfully activate the thin film with magnetic fields less than 1 mT in strength, which is at least an order of magnitude smaller than other magnetically driven neural stimulation technologies

[16] . As a result, the inventors anticipate powering the Mag-Dust by small, lightweight magnetic field generators. These magnetic field drivers are also used to transmit data to the Mag-Dust ASIC (see the section entitled "Adaptive Data Reception").

[0155] Exemplary embodiments may include a wearable power and data transmitter including a battery, microcontroller, and H-bridge driver located external to the wearable pack. In certain embodiments, wires extend from the pack to a transmission coil located at the site of the implant. For example, as shown in panels (a)-(c) of FIG. 31, the coil may be located near the waist for a spinal implant, in the neckband for a vagus nerve implant, or in the visor around the head for a brain implant. By using a resonant circuit to lower the impedance of the coil, exemplary embodiments may create a 1 mT magnetic field with only a 0.2 mA draw on the battery, as shown in panel (d) of FIG. 31. With a 1200 mAh battery weighing less than 120 g, the battery life is estimated to be over 10 hours, even when operating continuously at full power.

[0156] Downlink data transmission to the ASIC can be achieved by modulating the driver frequency between resonant and off-resonant frequencies. When fully at resonance, the input voltage to the IC is at its maximum value. By selecting two alternative driver frequencies slightly off-resonant, the IC detects the changes in voltage generated by the membrane as information bits "0" and "1". In an exemplary embodiment, circuitry within the IC dynamically tracks these levels to compensate for coupling variations and driver amplitude drift, allowing for robust interpretation of voltage shifts as input data. Because the driver's frequency generator can generate all three frequencies simultaneously, frequency switching can be achieved seamlessly between stimulation pulses, minimizing dead time and maximizing data transmission rates. The precision of the frequency generation allows this approach to work reliably, even when the coil and membrane are designed to have a high Q value to maximize efficiency.

[0157] In vitro testing The functionality of the exemplary embodiment can be tested with ME-based wireless power and data transfer by attaching an LED to the stimulus driver and monitoring the optical signal using a camera or coupled optical fiber (see FIG. 32). This can allow remote sensing of the IC's response to power and programming without interacting with the magnetic field that drives the device. By measuring the LED's intensity and its output frequency and pulse width, we can verify both the reliability of our programming technique and the effective efficiency of the system across a myriad of test conditions. Test equipment can be used to examine the effects of distance and alignment between the coil and the device on performance and reliability. Such testing can be performed at 37 degrees Celsius in phosphate buffer solution to approximate biological conditions in the body.

[0158] example Modulation of electrical activity of the nervous system has shown great potential for neuroscience research and clinical therapy. To reduce the risk of infection and the restriction of subject mobility, neuromodulators should be miniaturized and cord-free. Safe and reliable wireless power transmission and data delivery, with size and power constraints, remain one of the fundamental challenges in developing miniaturized neural interfaces. Although several wireless neural implants powered by RF, inductive coupling [1-3], ultrasound [4] and light [5] have been reported; existing solutions cannot simultaneously achieve implant miniaturization, system portability, functional flexibility and subject mobility while avoiding tissue heating due to body absorption of high-frequency EM waves [6], attenuation of mechanical waves when transmitting through different media [7], limited penetration depth of light [5] or lack of controllability of magneto-thermal nanoparticles [8]. In comparison, magnetoelectric (ME) transducers (see top diagram in Figure 33), which convert low-frequency (100 kHz-10 MHz) AC magnetic fields into electrical energy via mechanical coupling between magnetostrictive and piezoelectric thin films, are promising for powering biomedical implants by offering the desirable properties summarized in Figure 33. At resonance, lead zirconate titanate (PZT) / Metglas-based ME thin films generate high output voltages (>7 Vpp) with low resistive source impedance (~800 Ω) under magnetic fields below 1 mT, making them suitable for efficient energy harvesting.

[0159] An exemplary embodiment of the present disclosure (sometimes referred to as the "MagMote") comprises a cordless, programmable neural implant that utilizes the ME effect. 2 The company integrates its 180 nm CMOS system-on-chip (SoC) with its in-house developed 4 mm × 2 mm ME thin film, a single energy storage capacitor, and on-board electrodes onto a flexible polyimide substrate. Specific embodiments are characterized by: (1) A 8.2 mm SoC that requires only two separate components to be added; 3and small physical dimensions of 28 mg; (2) adaptive system control and data transmission mechanisms robust under power supply amplitude variations (1 V variation tolerance and 0.2 V modulation depth); (3) 90% chip efficiency with low static power down to 23.7 μW; and (4) the ability to perform fully programmable biphasic current stimulation covering amplitudes of 0.05-1.5 mA, pulse widths of 64-512 μs and frequency ranges of 0-200 Hz, suitable for diverse clinical targets such as the spinal cord, vagus nerve and brain. The concept of a spinal cord stimulation system enabled by MagMote with a battery-powered magnetic transmitter assembled on a wearable belt is shown in Figure 33 (lower left panel).

[0160] As described herein, the MagMote SoC, including the power management, data recovery and stimulation modules, interfaces with the ME thin film to receive power and data and drive the programmable stimulation, as previously described in Figures 22A and 22B. The alternating current (AC) voltage induced on the ME transducer is first converted to a direct current (DC) voltage (Vrect) by an active rectifier. The energy is then stored in an off-chip 4.7μF capacitor. The LDO provides a constant 1V supply VLDO for clock and data recovery to reduce power consumption and effectively increase the input impedance of the SoC for a higher voltage conversion ratio. A temperature and supply invariant reference voltage for power management is generated by a native MOS-based reference circuit [9]. Downlink data is transmitted by amplitude shift keying (ASK) of the magnetic field and is used to program the stimulation current with 50μA resolution using a 5-bit current DAC. A finite state machine (FSM) configured by the received data also controls the timing of biphasic stimulation, which, unlike monophasic pulses, is charge balanced to prevent undesired electrochemical reactions. By the end of stimulation, the electrodes are shorted to completely remove any residual charge. The SoC cycles between a charging phase, a data reception phase, and a stimulation phase at the desired stimulation frequency set by an external transmitter (TX) (Figure 22, bottom left).

[0161] To maintain reliable system operation, especially in the presence of ME induced voltage amplitude variations due to distance changes and misalignment, the transitions between the three operational phases and digital communication are designed to be adaptive, calibration-free, and fully controllable by the TX (user) side. First, the enablement of the data transmission phases is controlled by varying the ME amplitude, and the time allocation for each phase is fully controlled by the external TX with precise timing reference and computing resources for control. To save power and area, the existing comparator in the active rectifier is reused as a watchdog to monitor the amplitude change. This comparator generates a pulse train that tracks Vrect during charging and stops if the amplitude of the input falls below the voltage of the energy capacitor. By comparing the watchdog signal with the divided Vrect and extracting the envelope, a rail-to-rail enable signal (ENdata) for data transmission is created (Figure 34, top right). Comparator offsets can also cause the tracking of Vrect to fail and generate erroneous ENdata (Figure 34, bottom left). To mitigate this, a four-input comparator

[10] , which has higher drive capability than a two-input comparator of the same size, increases the offset tolerance from 11 mV to 23 mV, effectively reducing the failure rate from 8% to 0.02% (estimated by simulated offset variance). Second, the system clock is derived directly from the ME source, and therefore presents a PVT-invariant frequency and strict duty cycle distribution (Figure 34, right). Third, to ensure correct data demodulation under ME voltage fluctuations, a voltage threshold is generated online before each data transmission using an alternating pilot tone sent by the TX. The threshold is extracted by a low-pass filter and then processed by a track-and-hold circuit. Measurement results show normal data transmission at 0.2 V modulation depth. Furthermore, we verified the robust operation of the device by showing that a fixed setting of stimulation current (maximum setting, bottom right of Figure 34) remained near the expected value of 1.5 mA as the ME-induced voltage varied from 2.6 to 3.6 V due to magnetic field fluctuations.

[0162] Figure 35 shows the measured waveforms when the MagMote is operated at 20, 67 and 200 Hz stimulation frequencies. rect The ASK reaches a peak of 2.95 V at 83% voltage transfer rate at 0.61 mT field strength and drops to 2.55 V after 1.5 mA, 512 μs stimulation. ASK is achieved by vibrating the ME film off-resonance at a reduced amplitude by slightly shifting the frequency. Set by the settling time of the ME film, 32 cycles are required to reliably transmit one bit, resulting in a data rate of 7.8 kbps. In addition, biphasic stimulation currents from 0.05 to 1.5 mA have been verified. The SoC consumes 23.7 μW of power without stimulation and achieves 90% peak efficiency at a maximum average stimulation power Pstim,avg (Pstim,avg=Istim2×Rload×Tpulse×fstim) of 2.27 mW.

[0163] To demonstrate suitability for implantation, the ME thin film is packaged in a 3D-printed enclosure to eliminate environmental effects on mechanical vibrations, and the entire implant is then encapsulated with non-conductive epoxy. Immersion testing is performed by immersing the MagMote in phosphate-buffered saline (PBS) for 7 days. At start-up, when centered on the TX coil, the device charges to 2.6 V within 44 ms with an average input power Pin,avg of 0.37 mW. At a position 30 mm away from the TX, a Vrect of 2.05 V and Pin,avg of 0.16 mW are achieved, which is sufficient to ensure SoC functionality. In addition to testing in PBS, the device is also tested in air and agar (a substrate used to mimic the mechanical properties of the brain) to show Vrect and Pin,avg variations of less than 0.14 V and 0.05 mW, respectively (bottom left of Figure 36), demonstrating the adaptability of the ME effect in different media. Throughout a one-week immersion test, the MagMote in PBS performed consistently, with Pin,avg varying between 0.35-0.39 mW, indicating long-term reliability for implantation.

[0164] To further evaluate the biostimulation capabilities of the MagMote, we validated the device using Hydra vulgaris as a model of excitable tissue. The transgenic Hydra strain used expresses the calcium-sensitive fluorescent protein GCaMP6s in the ectoderm and naturally expresses voltage-sensitive ion channels (Figure 37, top left). When applying biphasic pulses of 500us pulse width at 100Hz for 5 seconds, we observe a >200% increase in GCaMP6s fluorescence (Figure 37, top right), consistent with electrical activation of voltage-gated ion channels in Hydra tissue that produces muscle contraction. The table in Figure 37 compares the MagMote to state-of-the-art implantable wireless stimulators. A die micrograph of the MagMote SoC is included in Figure 38.

[0165] Figure 39 shows the experimental setup to capture fluorescence images using a microscope and constant excitation light; the stereotrode connected to the proposed device; and the laser used to image GCaMP6s activity in Hydra.

[0166] FIG. 40 shows a portable magnetic field driver and its system diagram.

[0167] All methods disclosed and claimed herein can be accomplished and carried out without undue experimentation in light of the present disclosure. Although the compositions and methods of the present invention have been described in terms of preferred embodiments, it will be apparent to those skilled in the art that modifications may be made to the methods and steps or the sequence of steps of the methods described herein without departing from the concept, spirit and scope of the invention. More specifically, it will be apparent that certain chemically and physiologically related agents may be substituted for the agents described herein with the same or similar results. All such similar substitutes and modifications apparent to those skilled in the art are deemed to be within the spirit, scope and concept of the invention as defined by the appended claims.

[0168] V. References The following references, to the extent that they provide exemplary procedural or other details supplementary to those set forth herein, are specifically incorporated herein by reference. TIFF0007637631000001.tif199159TIFF0007637631000002.tif243143TIFF0007637631000003.tif228156TIFF0007637631000004.tif241143

[0169] Further reading TIFF0007637631000005.tif160141

[0170] Further reading TIFF0007637631000006.tif162140

[0171] Further reading TIFF0007637631000007.tif203141

Claims

1. a magnetoelectric thin film including a central portion; a magnetic field generator configured to generate a magnetic field at an input frequency between 20 and 500 kHz; and an electrical circuit coupled to the central portion of the magnetoelectric thin film; Including, the magnetoelectric thin film configured to generate an electrical output signal when stimulated by the magnetic field at the input frequency; the electrical circuit includes an electrical component configured to modify the electrical output signal such that the electrical output signal has an output frequency of 0.5 to 1000 Hz; Device.

2. the magnetoelectric thin film has a resonant frequency; the input frequency is within 10% of the resonant frequency of the magnetoelectric thin film; 2. The apparatus of claim 1.

3. 10. The apparatus of claim 1, wherein the magnetic field generator is configured to generate a series of magnetic field pulses.

4. 4. The apparatus of claim 3, wherein each pulse in the series of pulses has a duration of 1 millisecond (ms).

5. the electrical output signal having an output voltage; The amplitude of the output voltage can be controlled by varying the duration of the pulses in the series of pulses.

4. The apparatus of claim 3.

6. the electrical output signal having an output voltage; The amplitude of the output voltage can be controlled by varying the total number of pulses in the series of pulses.

4. The apparatus of claim 3.

7. 10. The device of claim 1, wherein the electrical output signal is a monophasic output signal.

8. 10. The device of claim 1, wherein the electrical output signal is a biphasic output signal.

9. 2. The device of claim 1, wherein the magnetoelectric thin film comprises a magnetostrictive layer and a piezoelectric layer.

10. 2. The device of claim 1, wherein the magnetoelectric thin film comprises a polyvinylidene fluoride (PVDF) layer.

11. 2. The apparatus of claim 1, wherein the magnetoelectric thin film comprises a lead zirconate titanate (PZT) layer.

12. 10. The apparatus of claim 1, wherein the input frequency is between 20 and 500 kilohertz (kHz).

13. 13. The device of claim 1, wherein the output signal stimulates biological tissue.

14. 10. The apparatus of claim 1, wherein the output signal powers an electronic device.

15. 10. The device of claim 1, wherein the magnetoelectric film is a first magnetoelectric film, and the device further comprises a second magnetoelectric film.

16. 10. The apparatus of claim 1, wherein the magnetic field generator comprises an electromagnet.

17. 10. The apparatus of claim 1, wherein the magnetic field generator comprises a permanent magnet.

18. 10. The apparatus of claim 1, wherein the electrical components include a full-wave rectifier and a transistor.

19. 10. The apparatus of claim 1, wherein the electrical components include a diode and a capacitor.

20. 10. The apparatus of claim 1, further comprising a stereotrode, the electrical output signal being transmitted through the stereotrode.

21. the magnetoelectric film has a major axis; and A first end, a second end, a center point equidistant between the first end and the second end, and a length extending between the first end and the second end along the longitudinal axis of the membrane. Including; the central portion comprises 10% of the length along the major axis of the magnetoelectric thin film; the center point being located at the center of the central portion along the longitudinal axis of the magnetoelectric thin film; 2. The apparatus of claim 1.

22. 10. The apparatus of claim 1, wherein the magnetic field generator is configured to generate a magnetic field at an input frequency within 5% of a resonant frequency of the magnetoelectric thin film.

23. 1. A method of stimulating neural tissue in a non-human mammal, comprising: Providing a device according to any one of claims 1 to 22; generating a magnetic field with a magnetic field generator; generating an electrical output signal with the electromagnetic thin film; and modifying said electrical output signal with an electrical circuit. A method comprising:

24. 24. The method of claim 23, further comprising stimulating neural tissue with the altered electrical output signal.

25. 24. The method of claim 23, further comprising powering a device with the modified electrical output signal.

26. The device of claim 1 , wherein at least a portion of said device is implantable, said at least a portion of said device including said magnetoelectric thin film and said electrical circuitry.

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