JP2023106499A - midfield coupler - Google Patents

Abstract

To provide devices, systems and methods for use in wireless electric power transmission using midfield sources and implants.SOLUTION: Midfield sources can be implemented by a patterned metal plates constituted of one sub-wavelength structure among multiple structures. These midfield sources operate evanescent fields of the outside of substances (e.g., tissue), and excite and control propagation fields in the inside of the substances (e.g., tissue), and thereby can generate adaptive energy transfer being spatially shut in the substances (e.g., tissue). The energy can be received by an embedded device capable of being constituted for one or more functions such as stimulation, detection or drug delivery.SELECTED DRAWING: Figure 3

Description

(Cross reference to related applications)
This application claims the benefit of priority to U.S. Provisional Patent Application No. 61/994,960, entitled "MIDFIELD COUPLER," filed May 18, 2014, the entire contents of which are incorporated by reference. incorporated herein.

The present disclosure relates generally to wireless power transfer. More specifically, described herein are devices, systems, and methods for mid-field coupling to implantable devices such as microstimulators, sensors, ablation, drug delivery devices, and the like.

Despite considerable progress in energy storage technology, batteries remain a major obstacle to the miniaturization of implantable electronic devices. As a result, current implantable electrical stimulation systems typically include a large impulse generator containing a titanium case surrounding a battery and circuitry used to generate the electrical pulses. Impulse generators are typically implanted in body cavities such as under the collarbone, under the ribcage, in the lower abdominal region, or above the buttocks. Electrical pulses are then delivered to the target nerve or muscle area via leads under the skin or through blood vessels. Problems associated with this current approach include pocket infections, lead dislodgment, lead crushing or perforation, muscle tear resulting from lead implantation or withdrawal, and limited space for electrode placement. include. Additionally, the life of these devices is severely limited, requiring periodic surgical replacement once the battery unit is depleted.

Alternatively, energy can be wirelessly transmitted from an external source, but it remains difficult to be able to transmit power to small implanted devices and/or devices located beyond skin depth. Most of the known wireless powering methods for implantable electronic devices are based on near-field coupling methods, and these and other proposed methods have a number of drawbacks. Power harvesting structures (eg, coils or antennas) in implanted devices are typically large. The largest dimension is typically on the order of one centimeter or more. Coils outside the body in near-field coupling are also typically bulky and inflexible. This poses some difficulties in integrating external devices into daily life. The inherent exponential decay of the near-field limits the miniaturization of implanted devices beyond skin depth (greater than 1 cm). On the other hand, far-field radiation properties severely limit energy transfer efficiency. Accordingly, it may be desirable to have devices and methods for transmitting wireless power to small implantable devices and corresponding small implantable devices suitable for minimally invasive delivery methods.

Described herein are devices, systems, and methods for wireless power transfer utilizing midfield sources and implants. In one variant, the midfield source (coupler) can be realized by a patterned metal plate consisting of one or more sub-wavelength structures. These mid-field sources manipulate evanescent fields outside the material (e.g., tissue) to excite and control propagating fields inside the material, thereby generating spatially confined adaptive energy transport within the material. can be generated. Energy may be received by an implanted device that may be configured for one or more functions such as stimulation, ablation, sensing, or drug delivery, among other functions.

In one variation, the devices described herein are midfield sources. A midfield source may comprise a midfield plate and one or more excitation ports. A midfield plate may comprise a flat surface and one or more sub-wavelength structures. The midfield source may be configured for wireless power transfer through tissue. In some variations, the midfield source comprises a planar structure including at least one of a metal, a slot or metal strip, and an excitation port coupled to the slot or metal strip, the source being within the following equation: An electromagnetic field can be generated that has the spatial frequency spectrum of the field adjacent to the source with some non-negligible component.
In some of these variations the device comprises at least two slots or at least two metal strips. In some of these variations, two slots or two metal strips are excited by the same excitation port. In some of these variations, two slots or two metal strips are excited by the same excitation port using microstrip transmission lines. In some variations the device further comprises a controller for dynamically shifting the focal region of the electromagnetic field. In some variations, the device comprises eight slots arranged in pairs of intersecting straight and curved slots. In some of these variations each pair of slots is excited by a single excitation port.

Further described herein is a system for wirelessly powering an implant through tissue. In some variations, the system may comprise a midfield plate comprising a planar structure and a subwavelength structure, a midfield source comprising an excitation port for exciting the subwavelength structure, and an implant comprising a receive coil. , the mid-field source is configured to transmit power to the implant via tissue propagation modes. In some of these variations, the source may produce an electromagnetic field having a spatial frequency spectrum of the field adjacent to the source with a non-negligible component within Equation (1). In some of these variations, the midfield plate comprises a flexible substrate. In some of these variations, the flexible substrate includes an adhesive and is configured to adhere to the patient's skin. In some of these variations, the implant diameter is less than 3 mm. In some of these variations the implant comprises electrodes. In some variations the implant comprises a sensor. In some variations, the midfield source includes a controller that dynamically shifts the focal region of the electromagnetic field in response to feedback from the sensor.

Further described herein is a method for wirelessly transmitting power through matter to an implant. In some variations, the method includes generating an electromagnetic field by a source and wirelessly transmitting the energy through the material to a receiving coil of the implant, the spatial frequency spectrum of the field adjacent to the source being: contains non-negligible components in the range of Equation 1. In some variations, the source and implant are separated by at least 5 cm and the implant has a diameter of less than 3 mm. In some of these variations, the power transfer to the coil is at least 10 μW when 500 mW is coupled to the material. In some variations, the method further comprises wirelessly transmitting energy to a second receive coil of the second implant. In some variations, the method further comprises adjusting the focal region of the electromagnetic field.

Also provided herein are one or more sub-wavelength structures configured to wirelessly transmit power by manipulating an evanescent field outside the tissue to produce a spatially focused field within the tissue. and an implant configured to receive wireless power from an external module, the implant comprising at least one sensor or stimulator. In some variations the sensor is selected from the group consisting of thermal sensors, chemical sensors, pressure sensors, oxygen sensors, PH sensors, flow sensors, electrical sensors, strain sensors, magnetic sensors, and image sensors. In some variations the stimulator is selected from the group consisting of an electrical stimulator, an optical stimulator, a chemical stimulator, and a mechanical stimulator. In some variations, the implantable device includes a modular design that allows for interchangeable sensors and/or stimulators. In some variations, the one or more sub-wavelength structures are selected from the group consisting of patches, PIFAs, slots, cross-slots, aperture-coupled circular slots, and half-slots. In some variations the source is configured to adjust the position of the focal point of the spatially focused field. In some of these variations, the implant includes a sensor that detects the power level of the received wireless energy and provides feedback to an external module to automatically adjust the position of the focal point to optimize wireless power transfer. with a transmitter that converts In some variations, the implant is configured to be implanted on, in, or near the heart to apply leadless pacing to the heart. In some variations, the implant is configured to be implanted on, in, or near the brain to apply deep brain stimulation to the brain. In some variations, the implant is configured to be implanted on, in, or near the spinal cord to stimulate the spinal cord. In some variations, the implant is implanted over, within, or near the musculature of the tongue to stimulate the tongue to treat obstructive sleep apnea. Configured.

Further described herein is a method of cardiac pacing comprising the steps of implanting a wireless powered module in, on, or near the heart; transmitting a midfield propagating wave to the wireless powered module to power the module; A method is described that includes delivering, sensing a cardiac parameter using the module, and providing electrical pacing to the heart using the module.

Further provided herein is a method of deep brain stimulation comprising the steps of implanting a wireless powered module in, on, or near the brain; transmitting mid-field propagating waves to the wireless powered module to power the module; , sensing a parameter of the brain using the module, and stimulating the brain using the module.

Further provided herein is a method of stimulating tissue comprising the steps of: implanting a wireless powered module in tissue; transmitting a mid-field propagating wave to the wireless powered module to power the module; A method is described that includes sensing a tissue parameter and stimulating tissue with a module. In some variations, the method further includes adjusting the focus of the propagating wave to optimize wireless power transfer to the module. In some variations, transmitting includes transmitting a wave having a sub-wavelength structure that produces a magnetic field perpendicular to the wave and parallel to the tissue interface.

FIG. 4 is a schematic side view of a source positioned over the interface between air and high index material; FIG. 4 is a schematic diagram showing power transfer to a coil attached to the surface of the heart. FIG. 4 shows magnetic fields generated by optimized source current densities for power transfer through multilayer materials. FIG. 4 shows magnetic fields generated by optimized source current densities for power transfer through multilayer materials. 3A and 3C are diagrams showing magnetic fields generated by a mid-field source for power transfer through the multilayer material of FIG. 2B. 3B and 3D are diagrams showing magnetic fields generated by near-field sources for power transfer through the multilayer material of FIG. 2B. FIG. 4A illustrates the magnetic field produced by a mid-field source in the absence of the multi-layered material of FIG. 2B. FIG. 4B shows the magnetic field produced by a near-field source in the absence of the multi-layered material of FIG. 2B. 4 is a performance curve showing mid-field and near-field coupling for fixed power transfer efficiency; FIG. 10 is a top perspective view of a variation of a mid-field source described herein positioned near representative tissue; 1 is a schematic diagram of a sub-wavelength structure described herein; FIG. 1 is a schematic diagram of a sub-wavelength structure described herein; FIG. 1 is a schematic diagram of a sub-wavelength structure described herein; FIG. 1 is a schematic diagram of a midfield plate as described herein; FIG. 1 is a schematic diagram of a midfield plate as described herein; FIG. 1 is a schematic diagram of a sub-wavelength structure described herein; FIG. FIG. 3 illustrates the architecture of multiple pump ports; FIG. 3 illustrates the architecture of multiple pump ports; 1 is a schematic diagram of a midfield source with four sub-wavelength structures fed by one pump port; FIG. FIG. 2 is a schematic diagram of a mid-field source comprising an array of sub-wavelength structures approximating the optimal source; Fig. 2 is a schematic diagram of a mid-field source comprising an array of sub-wavelength structures approximating an optimal light source; FIG. 16A shows the performance of the midfield source of FIG. 15, and FIG. 16A shows the spatial frequency spectrum of the output field; FIG. 16B is a diagram illustrating the performance of the midfield source of FIG. 15, and FIG. 16B is a diagram illustrating power transfer efficiency; FIG. 16C shows the performance of the midfield source of FIG. 15 and FIG. 16C shows the scattering spectrum. 7B shows the spatial frequency spectrum of the output field from the sub-wavelength structure of FIG. 7A; FIG. 11B illustrates the magnetic field generated by the sub-wavelength structure of FIG. 11A; FIG. 11B shows the magnetic field generated by the sub-wavelength structure of FIG. 11B; FIG. 11C shows the magnetic field generated by the sub-wavelength structure of FIG. 11C; FIG. FIG. 11D illustrates the magnetic field generated by the sub-wavelength structure of FIG. 11D; 11E shows the magnetic field generated by the sub-wavelength structure of FIG. 11E; FIG. FIG. 1 shows a magnetic field generated by a conventional inductively coupled loop source; 16 shows a spatially-shifted focus field pattern produced by adjusting the relative amplitudes and phases between port signals in the mid-field source of FIG. 15; FIG. Fig. 3 shows the architecture of the midfield source controller; Fig. 3 shows the architecture of the midfield source controller; FIG. 10 illustrates the effect of real-time dynamic focusing; 1 is a photograph showing an example of an implant described herein resting on a human finger; FIG. 10 is a photograph showing an example of an implant described herein prior to epoxy encapsulation adjacent to a catheter sheath for size comparison; FIG. 1 is a photograph showing an example of an implant described herein inserted into the inferior epicardium of a rabbit. Figure 23C shows the operating frequency of a mid-field source applied to the implant of Figure 23C; FIG. 23C shows the ECG of the rabbit of FIG. 23C. Figure 23E shows the autocorrelation function of both off-resonance and resonance sections of the ECG signal of Figure 23E; 3 is a circuit diagram of a power receiving device; FIG. Fig. 3 is a circuit diagram of a data transceiver; Fig. 3 is a circuit diagram of the stimulator and sensor; 1 is a schematic diagram of an implant configured to stimulate tissue; FIG. 1 is a schematic diagram of an implant configured to stimulate tissue; FIG. 1 is a schematic diagram of an implant configured to stimulate tissue; FIG. FIG. 4 shows a magnetic resonance imaging reconstruction of an implant location within a porcine tissue volume. FIG. 4 shows a magnetic resonance imaging reconstruction of an implant location within a porcine tissue volume. Figures 27A and 27B show the specific absorption rate measurement setup. 27C and 27D show the specific absorption rate distribution from the setup of FIG. 27A. FIG. 27C shows power transfer measurements in the setup of FIG. 27A. FIG. 27D shows power transfer well below the threshold. Figure 27E shows the received power when the received power is coupled to the tissue in the setting of Figure 27B, and Figure 27F shows the received power is below the maximum permissible exposure level. FIG. 13 shows a midfield source applied to a patient; FIG. 2 illustrates a receiving device implanted within tissue; 1 illustrates a time domain multiplex communication system; FIG. Fig. 3 shows an amplitude and phase shift network; FIG. 12 shows a midfield coupler affixed to tissue; FIG. 10 illustrates a method of customizing a midfield coupler;

Problems with achieving wireless power transfer include the mismatch between the size of the implantable device/sensor and the power transfer source, the depth of the device/sensor within the patient, as well as the spatial placement of the device/sensor with respect to the power transfer source. can occur in Described herein are devices, systems, and methods for wireless powering of microimplants that may overcome these problems. The power sources described herein can generate non-stationary evanescent fields that can induce energy transmission through propagating modes of tissue volumes. The region of energy transmission, called the midfield of the electromagnetic field, is at a wavelength distance from the source that corresponds approximately to the wavelength of biological matter.

In conventional wireless power delivery approaches using near-field coupling (inductive coupling and its resonance-enhanced derivatives), the evanescent component outside the tissue (near the source) remains transient inside the tissue, leaving effective depth penetration does not allow Unlike conventional near-field coupling, the energy from mid-field sources is primarily transported in propagating modes, so that the transport depth is limited by environmental losses rather than near-field intrinsic exponential decay. Theoretical studies have shown that energy transfer implemented using these properties is two to three orders of magnitude more efficient than near-field systems. In mid-field coupling, tissue can act as a dielectric for tunneling energy, and coherent interference of propagating modes confines the field at the focal plane to dimensions much smaller than the vacuum wavelength, making it diffraction-limited for high-index materials. Spot size can be affected. By placing the implant in this high energy density region, the implant can be several orders of magnitude smaller and placed deeper into the material than systems using conventional wireless powering methods. In fact, because of the high dielectric constant exhibited by living tissue at microwave frequencies, the power sources described herein can deliver electromagnetic energy to implantable devices measuring one millimeter or less embedded deep within the body. can be configured to deliver

Theoretical Basis The underlying physics of the devices, systems and methods described herein arise from optimization techniques that limit the performance achievable by any physical implementation of the power supply. Power transfer occurs when the source interacts with the receive coil to produce a field that causes work extraction by the load in the receive circuit. In sub-wavelength receive coils, only the lowest order modes are of interest, and the power transfer mechanism can be explained by the electromagnetic induction properties of the dynamic magnetic field interaction. The power transmitted to the coil is
where B _S is the magnetic field produced by the source and M _C is the induced magnetization due to the current in the coil. The electric and magnetic fields produced by the time-harmonic current density _J on the surface of the source conductor can be determined by decomposing the current density into its spatial frequency components, each of which is represented by refraction and refraction across a planar interface. Follow the usual laws of reflection. Using the phasor representation with the time dependence of exp(-iωt), efficiency is defined as:
Formally, η is the ratio of the power available in the coil to the total power absorbed. Although this equation only considers dissipation in tissue and in practice other losses such as radiation and resistive losses may occur, the amount of power that can be coupled into the body is effectively the electric field Limited by induction heating.

This efficiency is field specific within the tissue multi-layer structure and represents an upper bound on the efficiencies that can be obtained. An expression for this efficiency can be derived using coupled mode theory. The exchange of energy between the source and receiver is given by the equation
indicated by where a _n is the _amplitude normalized so that _| ^a _n | , and κ is the coupling coefficient. It may be advantageous for the source and receiver to operate at resonance ω=ω _S =ω _C . The power transfer efficiency is defined as follows.
In the limit of weak coupling |κ| ² /Γ _S Γ _C <<1, the equation transforms to:
This is the product of two efficiencies. The coefficient on the left can be understood as the efficiency of power transfer to the coil in the absence of load. The coefficient on the right corresponds to the efficiency of power extraction by the load, and this coefficient is maximized when the impedance matching condition Γ _C = Γ _L is met. From standard power reasoning, the efficiency on the left can be shown as follows.
This is the efficiency defined above. Equivalent representations can be obtained using other forms for coupled electrical systems, such as a two-port lumped element network.

The source J _S can be chosen to maximize efficiency. A global optimal solution can be analytically solved for a particular power delivery configuration by defining the current in terms of the component tangential to the plane between the source structure and the tissue. For all sources, the electromagnetic equivalence theorem allows such two-dimensional current densities to be selected from the universal set S indistinguishable in the lower half-space of z<0 from the physical sources of the field. . Remarkably, a solution to the optimization problem that maximizes (8) can be found in closed form as a result of S's vector space structure. In contrast to local optimization algorithms, this resulting solution is a hard bound on the performance that can be achieved by any physical implementation of the wireless power supply.

This theory can be applied to systems with external power sources and implanted devices, as described herein. FIG. 1 is a schematic side view of source 102 positioned over the interface between air 104 and high index material 106 . Source 102 may produce an in-plane source current density J _S . This source current can produce an electric field E _S and a magnetic field H _S as shown by the dyadic Green functions G _E and G _H .
where ω is the angular frequency and μ ₀ is the air permeability. Applying a Fourier transform to each of the abscissas yields the spatial frequency spectrum of the field.
If k ₀ is the wavenumber of air, then the spectral content in air is
, which corresponds to the evanescent field in air.

Energy transfer from source 102 to a receive coil located within high-index material 106 can occur when source 102 interacts with the coil to produce a field that results in work extraction by a load within the receive circuit. The time-averaged power delivered to the coil can be expressed as:
where M _C is the induced magnetization on the coil. Therefore, the power transfer efficiency for the field is defined as follows.
where ε is the dielectric constant of the material and E _C is the electric field generated by M _C . Formally, η is the ratio of the power available in the coil to the total power dissipated in the material.

The power transfer efficiency defined by η above can be changed by the selection of the source J _S . If the air matter medium has a multi-layer structure, the global optimum can be solved analytically. By searching such global solutions over a range of frequencies using an appropriate dispersion model of biological material, the optimal power transfer for a particular biological tissue(s) can be determined. For example, the above theory can be used to determine the optimal power transfer for the approximation of the chest wall structure. FIG. 2A is a schematic diagram of power transfer to a coil attached to the surface of the heart. The chest wall may be approximated by a multi-layer structure as shown in Figure 2B. As shown in FIG. 2A, the power delivery configuration consists of a source 202 positioned on the skin (described in detail below) and a receive coil 204 (described in detail below) inserted into the heart tissue layer. can be configured. The optimal power transfer of this approximated chest wall structure can be determined by finding the optimal source J _S over a range of frequencies using a suitable dispersion model for a suitable biological material.

In this example, optimal power transfer may occur at 1.6 GHz. To determine this, the theoretical efficiency vs. frequency curve was oriented in the x and z directions with the optimum η in a multilayer model of tissue (1 cm space, 4 mm skin, 8 mm fat, 8 mm muscle, 16 mm bone, ∞ heart). were generated by searching over a wide search range (10 MHz to 4 GHz) of the coils that were used. The upper frequency limit was chosen to be about the self-resonant frequency of the coil. Coil losses were considered using an analytical model of a loop of wire embedded in uniform tissue and impedance matching by imposing the constraint Q<10, where Q is the quality factor. Using the Debye dispersion model for each tissue type, peak efficiency was found to occur at 1.6 GHz.

As shown in FIG. 2B, the J _S solution can produce a highly oscillating current density 206 that can cause the output field to focus on the receive coil 204. FIG. These fields were calculated from the spectral components of the in-plane source current density J _S (k _x , k _y ) using the dyadic Green function method. This method reduced to a simple transfer function because the plane wave component is the eigenfunction of propagation in multilayer structures. At each depth z, for example, to calculate the magnetic field H( _kx , _ky ,z)= _GH ( _kx , _ky ,z) _JC ( _kx , _ky ), the dyad _GH (k _x , k _y , z) were applied. An inverse Fourier transform yields a field for each depth.

FIG. 2C shows the spatial frequency spectra in the depth plane corresponding to the source (z=0), the skin surface (z=z _skin =−1 cm), and the coil (z=z _coil =−5 cm). . At the depth planes corresponding to the source and skin surfaces, the output fields 208, 210 may each be composed of significant evanescent components corresponding to the following equations.
Near the receive coil 204, the output field 212 may consist of significant propagating modes corresponding to the following equations.
where k _muscle is the wavenumber of the muscle tissue. Thus, due to the high dielectric constant exhibited by biological materials at microwave frequencies, perfect control of the propagation modes in tissue affected the evanescent wave component whose source "lens" is within the equation can sometimes be achieved.

FIG. 3A shows a time snapshot of the output magnetic field from a midfield source with such a lens, and FIG. 3C shows the normalized spatial frequency spectrum of the output field adjacent to the midfield source. FIG. 3B, on the other hand, shows a time snapshot of the output field from a corresponding near-field source with a coil having a diameter of 4 cm and operating at 10 MHz, where the maximum electric field in tissue is Normalized to be the same as for 3B. FIG. 3D shows the normalized spatial frequency spectrum of the output field adjacent to the near-field source. As shown in FIG. 3B, fields from near-field sources decay and do not propagate much more quickly than fields from mid-field sources. Similarly, the spatial frequency spectrum shown in FIG. 3C exhibits Equation 16, while the majority of the components shown in FIG. Do not propagate.

Unlike conventional near-field coupling, mid-field powering takes advantage of the high dielectric constant exhibited by biological materials at microwave frequencies to facilitate energy transport. Therefore, the benefit may not be seen if there is no tissue between the source and receive coils. For example, FIG. 4A shows the output magnetic field from a mid-field source when tissue material is removed, and FIG. 4B shows the output magnetic field from a near-field source when tissue material is removed. As shown, the magnetic field from the midfield source shown in FIG. 4A may decay more rapidly than in the presence of tissue material and not propagate, as shown in FIG. 3A. .

The difference between power transfer using near-field coupling and power transfer using mid-field coupling can also be illustrated by a performance curve of receive coil diameter versus power transfer range. FIG. 5 is a diagram showing performance curves for a fixed power transfer efficiency η=10 ⁻³ and a circuit load of 10Ω. To achieve this efficiency at a given wavelength, the coil size and operating depth are constrained to be below the performance curve. Curves can be generated by solving for values of optimal power transfer in the air-muscle half-space for the indicated operating depths and tissue wavelengths.
For example, if the operating depth (distance between the source and the receive coil) is 5 cm, the near-field source may require the receive coil diameter to be at least 15 mm, while the mid-field source should have a receive coil diameter of at least 2 mm. Just This performance curve shows that it is possible to deliver electromagnetic energy deep within the body to implantable devices that are one millimeter or less in size.

Devices Described herein are devices, systems and methods for implementing a midfield source as theoretically described above. Realization of a source with the required lens properties as described above requires more complex electromagnetic structures than conventional coils or dipole elements. In one variation, the midfield source may be realized by a midfield plate, which may comprise one or more subwavelength structures, and an excitation port configured to excite the subwavelength structures. These mid-field sources generate evanescent fields near the source to excite and control propagating fields within a material (e.g., tissue), thereby producing a spatially focused adaptive steering field within the material. can be manipulated. Energy may be received by an implanted device, which is described in more detail below.

The systems described herein may enable wireless power transfer to implanted devices at depths unattainable with conventional inductive coupling techniques. In addition, the implant can be much smaller (e.g., one, two, or three orders of magnitude smaller) than the external power source and much shallower (e.g., one order of magnitude smaller) than the depth within the material (e.g., tissue). , two orders of magnitude, or three orders of magnitude smaller). Additionally, the power that can be transmitted to the implant via the systems described herein can be sufficient to power stimulation delivery and/or complex electronics within the implant.

External module (midfield source)
Described herein are mid-field sources configured to generate power that can be wirelessly transmitted to an implant. In some variations, the entire midfield source may be integrated into the handheld device. Therefore, midfield sources may be suitable for on-demand use. In other variations, the midfield source may be configured to be worn on the body or applied to the skin surface. 6A-6C are perspective views of variations of mid-field source 602 positioned near representative tissue 604. FIG. As shown there, each midfield source 602 may comprise a midfield plate that may comprise a planar surface 606 and one or more subwavelength structures 608 . One or more sub-wavelength structures 608 on the midfield plate may be excited by one or more radio frequency ports that may collectively form a midfield source, as described in more detail below.

Planar Surfaces In some variations, the planar surfaces of the midfield sources described herein may comprise solid substrates or plates. For example, the planar surface may, in some variations, include a glass epoxy laminate such as FR-4, which may include a patterned feed copper layer. In other variations, the planar surface may include Rogers or ceramic for lower substrate loss. A planar surface may have a generally planar shape and may have any suitable dimensions. The thickness may depend on the number of metal layers in the substrate and may range from about 1 mm to about 3 cm. In some variations, the solid substrate may be approximately 6 cm by 6 cm and have a thickness of approximately 1.6 mm.

In other variations, the planar surface may comprise a flexible substrate. In some variations where the planar surface comprises a flexible substrate, the flexible substrate may be an ultra-thin flexible substrate and may be configured to conform to uneven or curved surfaces such as a patient's skin. For example, the planar surface may comprise ultra-thin FR-4 in some variations. In some variations, a flexible substrate can have a thickness of about 10 μm to about 1 mm. More specifically, in some variations the flexible substrate may have a thickness of about 100 μm. The thickness may depend on the number of metal layers in the substrate and the separation between different layers.

As shown in Figures 28A-28C, the planar surface can be configured to adhere to the patient's skin. A spacer may be placed between the planar surface and the skin for insulation. In some variations, planar surface 2802 may be bundled together with battery 2804 and circuitry 2806 into a thin patch. This can be adhered to the skin as shown in Figure 28A. In another variation, as shown in FIG. 28B, the battery 2804 may be separate from the patch, and the patch and battery 2804 may be configured to adhere separately to the skin. In yet another variation, the battery 2804 and circuitry 2806 may be combined and configured to adhere to the skin separately from the patch, as shown in FIG. 28C. A flexible substrate configured to adhere to a patient's skin may be configured to be worn for any suitable period of time and may be configured according to the application. For example, in some variations, the systems described herein may be used for on-demand stimulation, where the planar surface may be used for a period of stimulation (eg, about 1 hour to about several hours). It should be understood that the planar surface may be adhered to the patient's skin for a longer period of time, although it may be left on the patient's skin for a period of time. In other variations, the systems described herein can be used to charge batteries within implanted devices. It is understood that in some such variations, the planar surface may be left on the patient's skin for between about 1 minute and about 10 minutes, although the planar surface may be adhered to the patient's skin for a longer period of time. want to be

Subwavelength Structures As mentioned briefly above, a planar surface may be combined with one or more subwavelength structures to form a midfield plate. A "sub-wavelength structure" may be defined with respect to the wavelength of the field. Any source structure with a wavelength dimension much smaller than the wavelength in air, λ ₀ , can be called a sub-wavelength structure, where λ ₀ is the wavelength in air and λ _material is the wavelength in the high-dielectric constant material. can. If the dielectric constant of the high-k material is n, then the wavelength in the high-k material is less than the root n (square root of n) times the wavelength in air, i.e., λ _material =λ ₀ /root n (square root of n). is. For example, the dielectric constant of muscle at 1.6 GHz is 54, so λ _material =λ ₀ /7.3. Therefore, any source structure with dimensions such as λ _material can be a sub-wavelength structure. More specifically, the maximum dimension of each sub-wavelength structure d may be between 0.1λ _material and 2λ _material . In this case, the sub-wavelength structure can generate an evanescent field, and when the sub-wavelength structure is placed in close proximity to the high-dielectric constant material, the evanescent field can induce energy transmission through the propagating modes of the high-dielectric material. . The spatial frequency spectrum of the output field adjacent to the source has (16) significant components, as explained in more detail above.

Sub-wavelength structures may have any suitable design configured to generate and manipulate propagating fields within matter (eg, tissue), as described in more detail herein. In some variations, the sub-wavelength structure may comprise slots in the ground plane. In other variations, the sub-wavelength structure may comprise metal strips or patches that may be placed over the substrate and may have a ground plane underneath, although a ground plane is not required. Metal strips or patches may comprise any suitable material such as (but not limited to) copper. Metal strips and metal patches can have any suitable thickness, such as (but not limited to) about 30 μm.

Examples of suitable sub-wavelength structures are shown in FIGS. 7A and 7B. In the variation shown in Figure 7A, the sub-wavelength structure may comprise two linear metal strips 702, 704 arranged end-to-end. The metal strips 702, 704 may be energized by a voltage source 706 (discussed in more detail below) located between their ends. The total length of the metal strips should be between about 1/10 the wavelength of the magnetic field in the dielectric material (e.g. tissue) and about twice the wavelength of the magnetic field in the dielectric material (e.g. tissue), as described above. obtain. A metal strip can be placed on top of the planar substrate as described above. In a variation shown in FIG. 7B, the sub-wavelength structure may comprise linear slots 710. FIG. The slots may be placed on the planar surface described above, such as metal plate 708 as shown, and slots 710 are energized by voltage source 712 to form a midfield plate (described in more detail below). can.

Other variations of sub-wavelength structures are shown in FIGS. 8A-8F. As shown in FIG. 8A, in some variations the sub-wavelength structure may comprise two curved metal strips 802, 804 that may form an arc. As shown in FIG. 8B, in some variations the sub-wavelength structure may extend around a complete arc to form a ring-shaped metal strip 806 . In other variations, the sub-wavelength structure may comprise one or more slots, as shown in Figures 8C-8F. In a variation shown in FIG. 8C, the sub-wavelength structure may comprise slots 808 forming arcs. In a variation shown in FIG. 8D, the sub-wavelength structure may comprise ring-shaped slots 810. FIG. In a variation shown in Figure 8E, the sub-wavelength structure may comprise two linear slots 812, 814 forming a cross. In the variation shown in FIG. 8F, the sub-wavelength structure may comprise linear slots 816 and curved slots 818 that intersect at their midpoints. In each variation shown in FIGS. 8A-8F, each sub-wavelength structure can be excited by a single voltage source 820 (described in more detail below). That is, there is only one point in the structure where the voltage is fixed.

Each of the variations of the sub-wavelength structures described herein can produce electromagnetic fields adjacent to the source, and the spatial frequency spectrum of these fields, as explained in more detail above, can be a few It has 16 non-negligible components. For example, FIG. 17 shows the spatial frequency spectrum of the transverse electric field adjacent to the midfield source with the sub-wavelength structure shown in FIG. 7A, where the total length of the metal strips in FIG. Approximately equal to λ _muscle at the operating frequency, the source is placed approximately 1 cm above the air-muscle interface. As shown in FIG. 17, the electric field contains non-negligible components of Eq.

In some variations, the sub-wavelength structure may be configured to minimize tissue heating effects of the applied field. Since electric fields induce tissue heating, sub-wavelength structures can be configured to produce a predominant magnetic field near the source in order to minimize tissue heating effects. Additionally or alternatively, the sub-wavelength structure can be configured to be thin. For example, sub-wavelength structures may be desirable with slots and/or patches because they are thin structures.

In other variations, the sub-wavelength structure can be configured to produce a dominant transverse magnetic field near the source. In some of these variations, the sub-wavelength structure is a patch sub-wavelength structure, i.e. a sub-wavelength metal plate on a substrate with a ground plane underneath (shown in Figure 11A), one side of the patch shorted to the ground plane. A PIFA sub-wavelength structure similar to the patch sub-wavelength structure (shown in FIG. 11B) except that it can be a cross-slot sub-wavelength structure (shown in FIG. 11C), the slot structure is close to the slot can include aperture-coupled circular slot sub-wavelength structures (shown in FIG. 11D), and/or half-slot sub-wavelength structures (shown in FIG. 11E), excited by monopoles that do not contact the slots.

18A-18E are diagrams showing the magnetic fields generated by the mid-field sources comprising the sub-wavelength structures of FIGS. 11A-11E. FIG. 18A is a diagram showing the magnetic field generated in a patch sub-wavelength structure such as that shown in FIG. 11A. FIG. 18B is a diagram showing the magnetic field generated in a PIFA sub-wavelength structure as shown in FIG. 11B. FIG. 18C shows the magnetic field generated in an aperture-coupled circular slot sub-wavelength structure as shown in FIG. 11D. FIG. 18D shows the magnetic field generated in a cross-slot sub-wavelength structure as shown in FIG. 11C. FIG. 18E shows the magnetic field resulting from a half-slot sub-wavelength structure as shown in FIG. 11E. As shown, the mid-field source produces a magnetic field parallel to the tissue interface and perpendicular to the propagating waves generated within the tissue that transmit wireless power to the implanted device. FIG. 19, on the other hand, shows the magnetic field produced by a conventional inductively coupled loop source. As shown, the magnetic field is generated perpendicular to the tissue interface and parallel to the direction of desired wireless power transfer to the implant placed in the tissue below the loop source.

Midfield Plate A planar surface as described herein can be combined with one or more sub-wavelength structures to form a midfield plate. The midfield plate may comprise any suitable number of sub-wavelength structures (eg, 1, 2, 3, 4, 5, 6, 7, 8, or more). Each of the sub-wavelength structures may be identical sub-wavelength structures, or the midfield plate may comprise a combination of various sub-wavelength structures as described above.

Figures 9A and 9B show examples of midfield plates that can comprise two or more sub-wavelength structures. In the variation shown in Figure 9A, the midfield plate may comprise two semi-circular sub-wavelength metal strips 902, 904 configured to form a complete ring-shaped structure. Two voltage sources 906 may excite the sub-wavelength structure. The configuration shown in FIG. 9A is similar to the configuration shown in FIG. 8B, but the configuration of FIG. 9A can have two voltage sources forming two points at which the voltage can be fixed. Accordingly, the configuration shown in FIG. 9A may comprise two sub-wavelength structures 902,904. FIG. 9B includes a cross shape similar to the configuration shown in FIG. 8F, but with four voltage sources 908 and four points at which the voltage can be fixed, thus four sub-wavelength structures 910, 912, 914. , 916 are shown.

In some variations, the midfield plate may comprise a combination of multiple sub-wavelength structures that approximate optimal source performance to maximize power transfer efficiency, as described above. Figures 14 and 15 show the configuration of such two midfield sources. In the variation shown in FIG. 15, the midfield plate may include an array of four configurations (discussed in more detail below) of sub-wavelength structures shown in FIG. 13A. The four configurations in FIG. 13A are arranged in a circular configuration with each of the four configurations rotated 90 degrees relative to the adjacent configuration, with the linear slots pointing toward the center of the array. A midplate with sub-wavelength structures in this arrangement, when excited, can produce a circular current path that can approximate the optimal current density J _S for powering the heart across the chest wall, as described above. In the variation shown, the midfield plate is excited to form the midfield source by four independent radio frequency ports connected to microstrip transmission lines, as described in more detail with respect to FIGS. 13A-13B. can. The amplitude and phase at each port can be selected to maximize power transfer efficiency. For proper phasing between port signals, the array structure can produce circular current paths that can approximate optimum current densities.

If the use of the midfield source shown in FIG. 15 is simulated through a chest wall approximated as shown in the arrangement of FIG. 2A, the spatial frequency spectrum of the array along the k _x axis is , shown in FIG. 16A compared to the theoretical optimum shown at 210 (skin surface) in FIG. 2C. As shown in FIG. 16A, the evanescent spectrum can be close to the theoretical optimum, but the radiation mode contribution can be about twice as large due to the inherent directivity of planar structures. Experimental measurements of power transfer efficiency are shown in FIG. 16B. Experimental studies have shown that when 500 mW of power is transmitted to an implant with a 2 mm diameter coil and immersed in a solution with dielectric properties that mimic muscle tissue, it has a sub-wavelength structure as shown in FIG. The mid-field source was able to obtain efficiencies within 10% of the theoretical coupling, as shown in FIG. Proven by value. It should be appreciated that the midfield plate of FIG. 15 can be modified in any way. For example, FIG. 10B shows a midfield plate with a configuration of sub-wavelength elements similar to the midfield plate of FIG. 15, except that the slots may include bends at each end. These bends can increase the bandwidth of the midfield source produced by exciting the midfield plate.

Planar Immersion Lens A solid immersion lens comprises a hemispherical dome of high index material placed on or near the air-matter interface, the hemispherical dome permitting light to reach these "forbidden" angles of refraction. enable This feature allows the light to be focused to a spot much smaller than the free-space wavelength, with a diffraction-limited resolution set by the matter wavelength ~λ/n. Solid immersion lenses are used extensively in many applications, including imaging, data storage, and lithography. However, solid immersion lenses are inherently three-dimensional and bulky. Replacing conventional solid immersion lenses with their flat counterparts offers opportunities for integration into complex systems, including nanophotonic chips, or biomedical devices that are conformal in the low-frequency regime.

FIG. 14 shows a planar immersion lens based on a metasurface. Metasurfaces are flat devices consisting of structural arrays of sub-wavelength apertures or scatterers that produce abrupt changes in electromagnetic properties as light propagates across the surface. The properties of the metasurface can be tuned by varying the parameters of the individual sub-wavelength elements to produce desired spatially varying responses. This design freedom is used to create devices that produce negative refraction, optical vortices, flat lens effects, holograms, and other special interface phenomena in both the optical and microwave domains. It's here. The planar immersion lens of FIG. 14 uses metasurfaces based on electrically thin metal strips with deep sub-wavelength spacing that allow radiation incident from air to be refracted to the “forbidden” angles of the material. This feature allows the creation of thin, flat devices that mimic the functionality of solid immersion lenses. This device can be fabricated on a flexible substrate and can operate at microwave frequencies.

A solid immersion lens is an optical tool that allows light entering matter from air or vacuum to be focused to a spot much smaller than the free space wavelength. Conventionally, however, solid immersion lenses rely on hemispherical topography, are non-planar and bulky, which limits their integration in many applications. A planar immersion lens is shown in FIG. The resulting planar device, when placed near the interface between air and a dielectric material, can focus electromagnetic radiation incident from air to a spot within the material that is smaller than the free-space wavelength.

Refraction at the air-matter interface determines the diffraction limit when light is focused from air onto a material. A conventional optical lens placed in the far field of the interface controls only the propagating wave component in air. As a result, their focusing resolution in matter is diffraction limited at the free-space wavelength λ. This is at least partially due to the inability of far-field light to reach higher wavevector components in matter. These high wavevector components correspond to plane waves propagating at angles greater than the critical angle and are confined within the material by total internal reflection.

To enable interfacial phenomena different from conventional reflection and refraction, metasurfaces can be placed at or near the air-matter interface to break the translational symmetry at the interface. A metasurface can impart a phase with a constant gradient ∇Φ to incident light, and propagation is governed by a generalized form of Snell's law. This law states that radiation incident at an angle θ _inc has a sufficiently large phase gradient, |∇Φ|>k ₀ −k ₀ sin |θ _inc |, if the forbidden angle |θ _ref |>θ _critical means to bend. A phase gradient can be realized by aperiodically modulating the surface with sub-wavelength structures of varying impedance. FIG. 14 shows the metasurface. A metasurface may include metal strips with passive lumped elements (resistors, capacitors, and inductors). At microwave frequencies, these elements can consist of patterned metal traces or commercial impedance elements. During resonance, the phase of the current in the strip differs from the phase of the driving electric field by a value between [0, π]. By choosing appropriate passive elements and considering both the intrinsic and mutual impedances of the structure, the spatial phase profile of the transmitted wave can be shaped within this phase range. The use of discrete passive elements greatly simplifies the design as the metasurface can be reconfigured simply by changing the elements. Since coupling is explicitly considered in the design, the spacing between elements can be sub-wavelength. The phase range can be extended to full [0,2π] by taking advantage of changes in polarization (Berry phase), incorporating elements with magnetic response, and/or cascading multiple layers, although a single layer The limited range achieved with the immersion lens of FIG. 14 using is also sufficient.

Refraction at the "forbidden" angle can be achieved using a metasurface to produce a phase gradient of ∇Φ=π/0.55λ for radiation at about 1.5 GHz. The spacing between elements is approximately λ/20, sub-wavelength to meet the sampling requirements. For an s-polarized plane wave incident at θ _inc =30°, the beam is completely refracted to an extraordinary angle of θ _ref =45° which is well above the critical angle θ _critical =30°. Since the metasurface does not rely on polarization conversion, there are no co-polarization components refracted at the angles prescribed by the standard Snell's law. Unlike diffraction gratings, which alter their spatial amplitude profile, metasurfaces refract the incident wave by modulating its phase profile, thus not producing unwanted diffraction orders.

Extraordinary refraction still occurs when a λ/20 air gap is introduced between the metasurface and the interface. This effect is closely related to unsatisfactory total internal reflection. In the absence of matter beneath the metasurface, the incident wave is completely reflected from the metasurface, forming an evanescent wave at the surface. This evanescent wave propagates along the surface in the direction of the phase gradient, a behavior that cannot be achieved with a grating. When a material is placed in close proximity to a metasurface, the phase of the evanescent field matches the propagating wave within the material, allowing the incident beam to tunnel into the material with a net transport of energy across the interface. . By varying the angle of incidence for the phase gradient, the angular spectrum of the transmitted beam can be placed almost entirely within the forbidden region. The result appears to be consistent with generalized Snell's law. The spread around the predicted angle is due, at least in part, to the finite size of the aperture.

To design a planar immersion lens, a field source in air focused to a λ/η spot in the material can be searched. Focusing across a planar interface has been investigated, but the conventional expression for an optimal field source considers only far-field light, yielding a ˜λ focal spot. A more general approach to describing evanescent waves at interfaces can be used. A spatial optimization problem of the current sheet j _s in the source plane (let z=0) is formulated. A solution may be defined to be the current sheet that maximizes the convergence metric. Assume that the material is dissipative and tolerates small but non-zero losses. The efficiency of work performed on matter as a focus metric is:
where _r is the focal point, the double prime of α is the imaginary part of the polarizability of the object at the focal point, the double prime of ε is the imaginary part of the permittivity of the material, and α is the focal point centered It can be set to the polarizability of a "virtual" sphere. The sphere has the same dielectric constant as the background material and can be made arbitrarily small (eg, diameter in computational mesh units). The electric field E is found by propagation from the current sheet js as described by the Green's function G(r,r').

Optimization problems can now be considered in terms of operators. Using Dirac's bracket notation, E and j _s can be expressed as functions of the Hilbert space |Ψ> and |φ>, respectively. They are related by the operator formula (19), where the hat of G is Green's functional operator. The hat of the focus position operator Φ can be defined such that the numerator of Equation 1 can be expressed as Equation 2 below.

Similarly, the hat of the power loss operator Σ can be defined such that the following result is obtained.

Optimal focusing occurs when the choice of source current density |φ> maximizes Eq. Convergence can therefore be presented as an optimization problem.
where S is the set of all current sheets on the plane above the z=0 plane. The form of Equation 4 is a generalized eigenvalue problem involving the operators (23) and (24).
The solution is obtained by a two-dimensional current density that satisfies Equation (25).
where λ _max is the maximum generalized eigenvalue. If the hat of B is invertible, then the solution |φ _max > is obtained from the standard eigenvalue decomposition of the operator (hat of B) ^-1 (hat of A). Numerical calculations are performed by (i) choosing a plane-wave basis that diagonalizes the Green's functional operator for the multilayer geometry and/or (ii) reducing the degeneracy due to the azimuthal symmetry about the focal axis. It can be greatly accelerated by using This computation reduces to the inverse of the binomial product at each spatial frequency without the need to form the full system matrix explicitly. This inverse filtering process is closely related to time reversal and can be generalized to transparent media by asymptotically driving the material loss to zero.

Consider a two-dimensional shape in which the material has a refractive index n=2. For incident S-polarized radiation, equation 4 can be solved numerically to obtain a line-focusing source at a distance of 4λ/n. A linear metasurface can be used to shape a normally incident plane wave such that the outgoing field matches the solution. The required impedance values of the passive components can be solved using a point matching method. The linewidth of the focal spot can be sub-wavelength, such as 0.42λ full width half maximum (FWHM). To verify that the focusing effect is due to phase (not amplitude) modulation of the incident wave, the passive element can be removed so that the surface acts as a grating aperture. The focal spot of the grating is not sub-wavelength (~λ), ie the intensity at the focal point is also reduced by a factor of four when the passive element is removed. The physics underlying the lens effect are substantially different from near-field focusing devices. Unlike near-field plates, which focus evanescent waves on strictly sub-wavelength intervals (typically less than λ/10), the focusing ability of immersion lenses arises at least in part from conventional interference between propagating waves. , so that the focal plane can have many wavelengths removed. The high-resolution enhancement of our lens results not from the near-field interference effects of the near-field plate, but from the shaping of the near-field phase profile, which couples the incident wave into forbidden angles when interacting with matter. Focusing is not dependent on near-field intrinsic attenuation, so the intensity of the focal spot can be comparable to or higher than the incident intensity. As with solid immersion lenses, focusing resolution remains diffraction-limited, but spot size is a function of matter rather than free-space wavelength.

Now consider a three-dimensional geometry in which a planar source is positioned a sub-wavelength distance (λ/15) above a material (real part n=8.8) whose refractive index at microwave frequencies approximates biological tissue. . Because of the symmetry about the focal axis, the polarization of the field at the focal point can be specified arbitrarily. If the electric field is set to be linearly polarized in the x-direction, the solution of Equation 4 can be a surface wave consisting of concentric circular currents about the focal axis. In air, the resulting field is an evanescent and non-stationary field, propagating in the plane towards the focal axis. The intensity profile in the source plane is significantly non-zero only within a finite circular region. The radius of this region defines the effective aperture size, which is directly related to the loss of the matter system and the depth of focus. At the focal plane of the material, the field converges to a spot of width λ/11 (measured at FWHM) at a distance of about 2.3λ/n (the wavelength of the material) from the source plane. Although the waves are generated in air, the spot size approaches the Abbe's diffraction limit λ/(2 nsin θ _ap ) for homogeneous materials. In this case, the source can reach the forbidden wave component, so θ _ap is the half angle at which the aperture corresponds to the focal point.

In short, FIG. 14 shows a midfield plate with the functionality of a solid immersion lens. The high focusing resolution of the device arises, at least in part, from the metasurface's ability to control the near-field with sub-wavelength resolution in its interactions with dielectric materials. At optical frequencies, a planar immersion lens can be realized by closely spaced plasmonic antennas or dielectric resonators whose interaction is accounted for by tuning the properties of the optical "focusing" elements. By incorporating sub-wavelength structures that interact with the magnetic field component of incident radiation, metasurfaces can further modify optical impedance to eliminate reflections at interfaces. Due to the inherently simple and planar fabrication of metasurfaces, metasurface-based lenses can be integrated into complex systems such as nanophotonic chips or conformal biomedical devices.

In the variation shown in FIG. 14, the midfield plate may comprise an array of curved sub-wavelength structures 1402, 1404, 1406, 1408, 1410, 1412, 1414, 1416 and sub-wavelength dipoles 1018. These can be nested to form an ox-eye pattern as shown. This choice of structure is configured to produce a propagating field converging to a sub-wavelength spot within tissue, as described in more detail above. In another variation, the midfield plate may comprise slots having the same configuration as shown in FIG. 14 instead of metal strips. 10A is a top view of the midfield source of FIG. 15. FIG. Since the midfield source of FIG. 14 has more excitation ports than the midfield source of FIG. 15, it can be more invariant to tissue properties between the source and the implant, whereas can work well with a wide variety of materials.

Excitation Ports The midfield plates described herein can be configured to manipulate the evanescent field produced by the power source. In some variations, the sub-wavelength structures of the midfield plate can be excited by excitation ports, as briefly described above. In some variations where the midfield plate comprises more than one sub-wavelength structure, each sub-wavelength structure can be excited by a separate excitation port. In other variations where the midfield plate comprises more than one sub-wavelength structure, a single excitation port can excite more than one sub-wavelength structure.

The excitation port may comprise a radio frequency port. A radio frequency signal may be generated by a signal generator (eg, an oscillator). A radio frequency signal may have any suitable frequency. In some variations, the frequency can be from about 800 MHz to about 1 GHz. In other variations, the frequency can be from about 2.3 GHz to about 2.5 GHz. As mentioned above, the optimum frequency for efficient power transfer can depend on the material placed between the midfield source and the receive coil. For example, in the example of a chest wall structure, the optimal frequency may be approximately 1.6 GHz. In some variations the frequency of the signal may be adjustable. By adjusting the operating frequency of the source, the power received by the implant can be adjusted and/or the mid-field source can be used for implants placed in different materials and at different locations within the material. .

In some variations with more than one excitation port, the radio frequency signal is split into multiple radio frequency signals, for example using a power splitter (e.g., Wilkinson power splitter) on the control board. obtain. In some variations, the radio frequency signal may be split symmetrically into each of a plurality of radio frequency signals, but this need not be the case. It should also be appreciated that rather than splitting a single radio frequency signal into multiple radio frequency signals, the device may comprise multiple signal generators. Each radio frequency signal can be transmitted from the control board to each radio frequency port via a cable (eg, semi-rigid coaxial cable). Additionally or alternatively, the signal is fed through a phase shifter with controllable phase (e.g. analog 400°, +3.5/-2.0° error) and/or subsequently amplified (e.g. , gain 14 dB). This allows controlled phase and amplitude signals to be generated at each radio frequency port.

Two exemplary architectures for multiple excitation ports are shown in FIGS. 12A and 12B. In the architecture shown in FIG. 12A, the signal generated by signal generator 1202 may be split, and each split signal is then fed through attenuator 1204, which may have a controllable variable attenuation setting. obtain. The signal may then be fed through phase shifter 1206 and amplifier 1208 . The architecture shown in FIG. 12B combines the amplifier and amplitude control element into a single component 1210 to produce the same controlled phase and amplitude signals with fewer components. can be done. Additionally, FIGS. 21A and 21B, described in more detail below, are another schematic diagram of a similar architecture for multiple pump ports 2110. FIG.

In some variations where the midfield source comprises two or more sub-wavelength structures, the two or more sub-wavelength structures can be excited by a single pump port. That is, the excitation for each sub-wavelength structure does not have to be done separately. This can be accomplished, for example, by carrying the signal from one radio frequency port to multiple sub-wavelength structures via microstrip. For example, FIGS. 13A and 13B show an example of a source with a sub-wavelength structure similar to that described with respect to FIG. 9B, with linear slots 1302 and curved slots 1304 intersecting about their midpoint. , where the four sub-wavelength structures described above are pumped using the same single pump port 1306 . A ring-shaped microstrip transmission line 1308 may be placed below a ground plane 1310 separated by a dielectric (eg, air or substrate 1312). Each point on microstrip transmission line 1308 may have a different phase. By adjusting the dimensions of microstrip transmission line 1308 , multiple sub-wavelength structures (eg, four sub-wavelength structures as shown) can be excited by pump port 1306 .

Although the signal generators described above are voltage sources, in other variations one or more sub-wavelength structures may be excited by current sources. In still other variations, the voltage or current sources may be replaced with reactive elements such as resistors, capacitors, inductors, or combinations of these elements. In these variations, the sub-wavelength structures can be excited by plane waves or waveguides. Therefore, the ratio of voltage to current can be fixed at the position of the reactive element instead of having a fixed current or fixed voltage.

Real-Time Dynamic Focusing In some variations of the midfield sources described herein, the focal region is controlled by the source's mechanical It can be dynamically shifted without reconfiguration. This may be due to the source being an implantable device configured to interact with an organ of rhythmic motion (e.g., for breathing or heartbeat), or to an implantable device configured to move within the body. It can be useful in clinical applications where it can be used to power. The excitation of individual sub-wavelength structures can be reconfigured in real time to shift the focal region, thereby synthesizing a variety of field patterns, including those with spatially shifted focal regions. can. FIG. 20 shows a field pattern with a spatially shifted focus designed by adjusting the relative phases between port signals. The upper diagram of FIG. 20 shows the formation of propagating waves in the direction just below the source. The bottom diagram of FIG. 20 shows focus adjustment.

Figures 21A and 21B show a possible architecture for a controller for a midfield source with four excitation ports. As shown in FIG. 21A, a radio frequency signal may be generated from a signal generator 2102 and symmetrically split into multiple radio frequency signals via a power divider 2104, such as a Wilkinson power divider. . Following power division, the signal may be connected in parallel stages for variable attenuation, phase shifting via element 2106 and amplification via element 2108 . This can produce controlled phase and amplitude signals at each excitation port 2110 . In other embodiments, following power division, the signal may be connected in parallel stages for phase shifting and variable amplification.

Phase shifters 2106 , 2108 may be controlled by control unit 2112 . In some variations, a "greedy" phase search algorithm may be used to change the phase and/or magnitude settings of each element of the midfield coupler to dynamically shift the focal region. In some variations, the algorithm may be based on closed-loop feedback, which may be relayed via optical fiber 2114, as shown in FIG. 21B. In other variations, the implant may be equipped with a radio transceiver that may be disconnected to implement closed-loop feedback and other related control algorithms. For example, feedback may be based on the detected power level of received wireless energy by the implant, as described in more detail below. Based on power measurement feedback from the implant module, adjustments can be made automatically and in real time to optimize wireless power transfer between the source and implant.

An example of the effect of such an adaptive algorithm is shown in Figures 22A-22E. To generate the image shown there, an implant 2216 with LEDs shown in FIG. ” moved in a trajectory. FIG. 22B shows the strobe position of the LED as the real-time control algorithm as described above is dynamically tracked for operation. FIG. 22C shows the strobe position of the LED without dynamic focusing. As can be seen by comparing Figures 22B and 22C, the field pattern is static and more centrally focused for the non-adaptive algorithm compared to the adaptive algorithm. On the 'S' shaped trajectory of motion, the adaptive algorithm has removed the dark areas that occurred in the static case, showing a much wider coverage area than is inherent in the focal area. Furthermore, the effect of the adaptive algorithm can also be seen in Figures 22D and 22E. FIG. 22D shows the power received by the implant as measured by the blink rate of the LED. As shown, the dynamic phase adaptation algorithm allows higher levels of power to be transmitted as the device moves. FIG. 22E shows the phase of each port relative to the phase locked port 4 controlled by the algorithm when the implant is moving.

Internal module (implants)
Also described herein are implants that may be configured to receive power from a midfield source as described herein. In some variations, the implant is configured to provide stimulation (e.g., electrical stimulation) to a target site (e.g., a targeted nerve, muscle, or tissue region), as described in more detail below. can be Additionally or alternatively, the implant may be configured to perform sensory functions at the target site, as described in more detail below.

A mid-field source may focus the output field to a sub-wavelength spot to produce a highly oscillating current density that may create a high energy density region deep in tissue. Inside this region, the power harvesting structures within the implant can be very small. Since the system can operate in the mid-field region, the implant can derive energy primarily from transverse electromagnetic fields (ie, the direction of oscillation of the field is perpendicular to the direction of propagation). This is in contrast to near-field coupling systems, where the implant can derive energy primarily from an axial electromagnetic field (ie, the direction of oscillation of the field is parallel to the direction of propagation). The focal region from the midfield source can be dynamically shifted using the degrees of freedom provided by the amplitude and phase of the input port signal, as described herein. The implant may incorporate components for received power sensing and wireless communication to enable real-time dynamic focusing as described herein.

Implants may consist of power receiving devices, data transceivers, and/or stimulation and sensing components. In some variations, the power receiving device may consist of a coil and one or more AC-DC conversion branches for different output voltage requirements. A data transceiver may consist of a data receiver, a data transmitter, a multi-access protocol and/or discriminator, and a digital controller. The stimulation and sensing components may comprise current drivers for both electrical and optical stimulation, sensing front ends for electrical sensing, electrodes, and/or LEDs (light emitting diodes). These components are described in more detail below, but it should be understood that an implant need not necessarily include all of these components.

Size and Shape Implants can have any suitable shape and dimensions. In some variations, the implants described herein may be configured to fit within a delivery device such as, but not limited to, a catheter or hypodermic needle. In these variations, the implant can be injected directly into the target site (eg, the targeted nerve or muscle area) without the need for leads and extensions. FIG. 23A is a photograph of an example implant 2302 described herein resting on a human finger. FIG. 23B shows the same implant 2302 prior to epoxy encapsulation adjacent to a 10 French (˜3.3 mm) catheter sheath for size comparison. FIG. 23C is a photograph of the same implant 2302 inserted into the inferior epicardium of a rabbit via open chest surgery. In the variation shown in Figures 23A-23C, the implant is a 2mm diameter wireless electrostimulator.

In some variations, the implant can have a cylindrical shape, semi-cylindrical shape, circular or rectangular shape, or the like. In some variations, the implant can have a diameter (or maximum cross-sectional dimension) of about 10 μm to about 20 mm, about 100 μm to about 10 mm, or about 1 mm to about 5 mm. More specifically, in some variations the implant can have a diameter (or maximum cross-sectional dimension) of about 2 mm. It should be appreciated that in other variations the diameter (or maximum cross-sectional dimension) of the implant may be greater than 20 mm. In some variations, the implant can have a height of about 10 μm to about 20 mm, about 100 μm to about 10 mm, or about 1 mm to about 5 mm. More specifically, in some variations the implant can have a length of about 3 mm. It should be appreciated that in other variations the length of the implant may be greater than 20mm.

In some variations, the implant may be encapsulated within a suitable substance. For example, in some variations the implant may be encapsulated in epoxy. In other variations, the implant may be encapsulated in ceramic or glass, and in some variations may include anchors or other structures to help secure the implant in place. In some variations, electrodes may be configured to be used for fixation in addition to stimulation and/or sensing. For example, the electrodes can include barbs, as shown in FIG. 23A. In other variations, the electrode can include a threaded feature that can be screwed into tissue. In other variations, the implant may comprise anchoring structures that are not electrodes. For example, the implant may include loops or hooks. Such implants may be secured to tissue by joining adjacent tissue via loops or hooks.

Coils Coils can be configured to receive energy from a source, such as a spatially adaptive electromagnetic field generated by a source described herein. Energy can be received by the coil as magnetization due to induced currents in the coil. The coil may comprise any suitable material such as, but not limited to, copper, gold, or aluminum. A coil may include any suitable number of turns. The number of turns may depend on the frequency of the midfield source. In some variations, the coil may include from about 1 to about 15 turns. FIG. 23B shows an example of a coil 2304 comprising a multi-turn coil structure comprising 200 μm diameter copper wire wound with an inner diameter of 2 mm.

AC-DC Conversion and Charge Pump In some variations of implants with an AC-DC power conversion mechanism, the AC-DC power conversion mechanism may include a rectifier circuit. The rectifying circuit is configured to convert the energy received by the implant (e.g., by the coils described above) (e.g., the spatially adaptive electromagnetic field generated by the sources described herein) into a DC signal. can be In some variations, the AC-DC conversion circuitry may be split into a low voltage domain and a high voltage domain. This can increase the efficiency of rectification and power management of wireless power implants operating in electromagnetically weakly coupled areas. In some variations the implant may comprise a charge pump. In one variation, two diodes (eg Schottky diodes) and two capacitors (eg 10 nF capacitors) can be placed in the charge pump configuration. At low frequencies, additional capacitors may be used to match the impedance of the coil and rectifier. A charge pump and flash control integrated circuit may be placed after the rectifier to upconvert the rectified voltage.

Integrated Circuits In some variations, the implants described herein may comprise integrated circuits. For example, as shown in FIG. 23B, coil 2304 can be disposed over and connected to integrated circuit 2306 . In some variations the integrated circuit may be configured to adjust the pulse amplitude. In some variations, the implant may comprise non-volatile memory. For example, the implant is configured to record data such as usage information (e.g., drive time and current driver settings) and/or store measurements from sensors (as described in more detail below). It may have a flash memory that can be In some variations, each implant module may have its own identification tag, such as an identification tag stored in the memory of the implant. In some variations, the implant can be configured to coordinate interactions between various components of the implant, communication between the implant and external components, and multi-access protocols, as described below. It can have a core.

Energy Storage Component In some variations, the implant may comprise an energy storage component. For example, an implant may include a rechargeable battery. Rechargeable batteries may be configured to temporarily store energy and/or be used as efficient charge pumps for power management circuits. In some variations, the rechargeable battery may comprise a thin film battery. In some of these variations, thin film batteries may be stacked to increase energy density. In some other variations, the rechargeable battery may include a lithium battery. The energy storage component allows the implant to operate without being continuously coupled to an external power source, as described herein. An external power source may be used to charge the implant, and in some variations the implant can be charged with just minutes to tens of minutes of wireless charging per week or month.

Sensors The power levels deliverable to implants by the power sources described herein can exceed the requirements of microelectronics technology, thus enabling real-time monitoring of chronic disease pathologies and closed-loop biological sensing and sensing by implants. More advanced functions such as controls may be implemented. In some variations, the implant may include one or more sensors. In some variations, the sensor may include, for example, a temperature sensor. In other variations, the sensors may comprise optical sensors and/or imagers. In still other variations, the implant may include chemical sensors, pressure sensors, oxygen sensors, pH sensors, flow sensors, electrical sensors, strain sensors, magnetic sensors, optical sensors, or image sensors. In some variations, sensors may allow the device to determine the depth at which it operates. In some variations, a sensor may comprise one or more electrodes.

In variations where the implant comprises one or more sensors, the implant may comprise one or more preamplifiers, analog/digital converters and/or drivers for one or more sensors. In a variant with an analog/digital converter, the analog/digital converter can be used to discretize the signal from the preamplifier. In some variations, the output signal from the analog-to-digital converter may be stored in the implant's non-volatile memory (as described in more detail above), or in other variations (described in more detail below). ) to a source or other external component via a radio frequency modulator.

In variations where the implant comprises one or more sensors, the sensors may be configured to provide feedback to the source or user. For example, in some variations an implant may include one or more sensors configured to detect instantaneous power levels received by the implant. This information may be transmitted to the source via a data transmitter (described in more detail below). This allows adaptive focusing of the focal region of the field, as described in more detail herein. In other variations, information from the sensors may be provided to the user, for example, via a user interface as described herein. In some variations, the data may be further analyzed or stored external to the implant. Information from one or more sensors, for example, enables wireless real-time monitoring, diagnosis, and/or treatment of a patient.

Stimulation In some variations, the implants described herein can be configured to deliver stimulation to tissue. Stimulation can be any suitable type of stimulation, such as, but not limited to, electrical, optical, chemical (e.g., implants can be configured for drug delivery), or mechanical stimulation. . In variations where the implant is configured to deliver electrical stimulation, the implant may deliver electrical stimulation via one or more electrodes. The implant may be equipped with programmable current drivers that allow it to deliver a range of stimulation parameters (eg, for electrical stimulation, intensity, duration, frequency, and shape). In some variations where the system includes a user interface as described herein, the programmable current driver can be programmed via the user interface (eg, via a wireless data link to the implant). In some variations, the implant may include two electrodes 2308, as shown in variations of FIGS. 23A-23C. Two electrodes 2308 may be placed on one side of the implant 2302 and inserted into tissue, as shown inserted into the inferior epicardium of a rabbit in FIG. 23C.

FIG. 24A is a circuit diagram of a power receiving device (implant). FIG. 24B is a circuit diagram of a data transceiver. FIG. 24C is a circuit diagram of the stimulator and sensor.

25A-25C are schematic diagrams of implants that may be configured to stimulate tissue (eg, implants configured as pacemakers). FIG. 25A shows a lumped circuit model of the receiver. As shown, an AC voltage Vc generated across coil 2502 by a source (as described in further detail herein) is converted to DC power via rectifier circuit 2504 . FIG. 25B shows the equivalent circuit at the nth reference power level when the circuit is in resonance. FIG. 25C is a diagram showing details of circuit components. As shown, the implant may include rectifiers, charge pumps, pulse control integrated circuits, storage capacitors, and LEDs. In another variation, the LED can be replaced with a pair of electrodes that can be configured for electrical stimulation. In some variations, the rectifier circuit may comprise two diodes and two capacitors arranged in a charge pump configuration. A charge pump can be placed after the rectifier to up-convert the rectified voltage required to drive the LEDs or electrodes, for example from 0.7V to 2V. Charge can be temporarily stored on a capacitor. A pulse control unit may be used to control the frequency and width of the pulses. In the example shown in Figures 25A-25C, the circuit further includes an LED configured to encode the power flowing through the pulse control circuit 2506 via the blinking frequency. The non-linear properties of the rectifier and pulse control unit allow the unknown parameter Rc to be estimated if the circuit is characterized with two reference blink frequencies. Knowing the environment dependent parameter Rc, the transmitted power can be estimated. In some variations, the LEDs of FIGS. 25A-25C may be replaced by a pair of electrodes that may be configured to stimulate tissue and/or nerves.

Data Transmission The implant is capable of wireless data transmission, and in some variations may include a wireless data link between the implant and a midfield source or another external component (eg, an external user interface). The wireless link may be unidirectional or bidirectional, so in some of these variations the implant may include a data receiver. The wireless link drives the implant (e.g., activates stimulation in variations with a stimulator), remotely programs or configures the implant (e.g., adjusts implant settings), and/or one or more can be configured to receive data from sensors of

The embedded module's data transmitter may use pulsed radio frequency modulation. In some variants, conventional load modulation may not work in midfield due to the low quality factor of implanted antennas, which can lead to poor signal-to-noise ratio and significant link margin variation Therefore, radio frequency modulation may be desirable. In other variations, non-coherent modulation techniques such as amplitude shift keying and frequency shift keying may be used for ease of implementation. The data carrier and power carrier may operate at different center frequencies to facilitate detection by an external module. In some variations, the implant may utilize a multi-access protocol that may modulate the functionality of the implant (eg, modulate multipoint stimulation). In some variations, the multiple access protocol may utilize time division multiplexing and frequency division multiplexing.

The data rate of the wireless data link may be any suitable rate (eg, several kbps to 10 Mbps). For example, in some variations the downlink data rate from the source to the implant may be a few Mbps or less, while the uplink data rate from the implant to the source is in the range of, for example, 10 Mbps or more. It can be a higher rate.

User Interface In some variations, the systems described herein may include a user interface that may be used by, for example, a clinician or patient. Some variations of the user interface may be integrated with the power supply, while in other variations the user interface may be separate from the power supply. In some variations where the user interface is separate from the power source, the user interface may comprise a mobile computing device such as a smart phone, tablet, or wearable computer. In these cases, the system may include a wireless or wired communication link that allows bi-directional communication between the power source and/or implant and the mobile computing device. This allows the patient and/or clinician to use the display of the mobile computing device to interface with the power source or implant (eg, receive or enter information).

Although the source, implant, and user interface are described herein as a system, it is understood that the devices described herein can be used alone or in combination with other devices and systems. want to be It should also be appreciated that the systems described herein may be configured based on the end-user's particular needs or requirements. For example, the implant may have a modular design and may be modified to include those components desired for the intended use. In some variations, all of the above components of the embedded module are multiple dies surrounded by a single die as a system-on-chip (SoC) or a single module as a system-in-package (SiP). can be integrated into

Methods/Uses Also described herein is a method of wirelessly powering an implant as described herein using a midfield source as described herein. there is The implants described herein can be implanted in any suitable location. In some variations, the implant may be implanted by a minimally invasive procedure such as a catheter or hypodermic needle. Implants can be implanted in humans or other animals such as pets, farm animals, or laboratory animals such as rabbits, mice, rats, or other rodents. Implants may be used for any number of purposes such as, but not limited to, muscle stimulation, stimulation/sensing to regulate a patient's heart rate, deep brain stimulation, drug delivery, and/or biological, physiological and chemical sensing. can be used for any purpose.

A mid-field source as described herein can be used to transmit power to the implant. In some variations, the power received by the implant may be adjusted by adjusting the operating frequency or other parameters of the midfield source, as described in more detail above. Additionally or alternatively, the parameters may be adjusted in real time to modify the focal region of the midfield source, eg, to track implant motion.

For example, in some variations the systems described herein may be used for cardiac pacing. In some of these examples, one implant may be delivered to the patient's right ventricle and a separate implant to the left ventricular epicardium. The implant may be delivered in any suitable manner, such as via a catheter through the vasculature (eg, the implant may be delivered to the left ventricular epicardium via the coronary sinus and coronary vein). These implants may include both stimulating and sensing electrodes that may be configured to apply leadless pacing to the heart. Thus, the system described herein allows leadless biventricular pacing to be achieved with only minimally invasive procedures. This can greatly reduce treatment time and complexity.

An example of wireless cardiac pacing using a system as described herein is shown in FIG. 23C described in more detail above. After implanting the wireless electrical stimulator in the inferior epicardium of the rabbit, the chest was closed, as shown in FIG. 23C. The electrical stimulator shown in FIG. 23C was approximately 2 mm in diameter, weighed approximately 70 mg, and was capable of producing 2.4 μJ pulses at a rate dependent on the extracted power. Its characteristic dimensions were at least an order of magnitude smaller than existing commercial pacemakers due to the lack of a battery. The portable, battery-powered midfield source shown in FIG. 29 was positioned approximately 4.5 cm above the implant 2302 (approximately 1 cm gap and 3.5 cm chest wall). Approximately 1 W of power was coupled into the chest using a mid-field source and the operating frequency of the source was adjusted to the estimated resonant frequency of the circuit and held for several seconds, as shown in FIG. 23D. Rabbit cardiac activity was monitored by ECG shown in FIG. 23E. The heart rhythm could be controlled completely wirelessly by adjusting the operating frequency. When the operating frequency matched the resonant frequency of the circuit, the pulse amplitude was sufficient to pace the heart, as indicated by the increased speed and regularity of the ECG signal in FIG. 23E. FIG. 23F shows the autocorrelation function for both off-resonance and resonance sections of the ECG signal. The peaks of the autocorrelation function are marked with squares.

Similar methods can be applied to any other photostimulation or electrical stimulation task in the body. For example, similar methods can be used to stimulate neurons or muscle cells. For example, similar systems and methods can be used for deep brain stimulation. Current procedures for deep brain stimulation involve drilling a hole greater than 1 cm in diameter in the skull and inserting leads and extensions from the leads to the stimulation module. Due to the invasiveness of the procedure, typically only a limited number of target sites are selected for electrode placement. Additionally, the leads may not be MRI compatible. Implants for use with the midfield sources described herein, on the other hand, can be injected into the brain via other minimally invasive routes and do not require leads or extensions. This allows more target sites to be stimulated and may be MRI safe. In addition, multiple devices can be implanted and used to stimulate synchronously. In addition, use of systems such as those described herein may reduce infection and reduce regulatory risk.

As another example, the systems and methods described herein can be used for spinal cord stimulation. Newer model spinal cord stimulator batteries are rechargeable because of their high power requirements. However, their power delivery approach relies on inductive coupling (or near-field coupling). Because the harvesting components are large in these systems, they can only be placed subcutaneously and not deeply. Therefore, the leads and extensions in these systems can potentially limit the position of the electrodes for effective stimulation. Lead dislodgment and infection can be a major cause of complications. Because the implants described herein are much smaller, the entire implant can be placed adjacent to the targeted nerve region of the spinal cord and may not require leads. Again, this may result in less infection, less damage to spinal cord tissue, and more effective stimulation.

As yet another example, the systems and methods described herein can be used for peripheral nerve stimulation. Most current devices support low-frequency stimulation and have much higher power requirements, so very few devices support high-frequency low-intensity stimulation. The system described herein can support both modes. Additionally, the two-way wireless link described herein may provide the ability to switch between different modes to tailor the stimulation paradigm to the individual patient.

As noted above, the systems and methods described herein can also be used to stimulate muscle cells. For example, the systems and methods described herein can be used to treat Obstructive Sleep Apnea (OSA). The implants described herein can be injected and implanted directly into the muscle tissue near the tongue and then used to deliver electrical stimulation to open the patient's airway during sleep. Multiple implant modules may be injected into different muscle groups to enhance muscle contraction. If desired, the patient can charge the implant with a midfield source. Additionally or alternatively, the data transfer function allows downloading the timestamp of each OSA episode that can be sent to the clinician. The implant can also be reprogrammed without being removed. In some cases, reprogramming may be based on collected data.

When the implants described herein are used to stimulate excitable cells, in some variations they may be used for temporary therapeutic applications where long-term implant implantation is undesirable. For example, currently, screening tests are typically performed before a permanent impulse generator is implanted. During screening exams, the patient can receive a temporary external impulse generator. The impulse generator may connect to extensions and leads that may be surgically placed within the body. During this period, the external impulse generator collects patient usage data and treatment efficacy. However, implants such as those described herein can obviate the need for a temporary generator with leads since they are injected into the target nerve/muscle area. Additionally, the implants described herein can be used to replace temporary sensing and pacing leads in patients after cardiac surgery.

The systems and methods described herein can be used for applications other than stimulation of excitable cells. For example, they can be used in medical sensing applications. Embedded sensors without a battery are typically passive sensors in nature, i.e., no active circuitry to condition the sensed signal exists within the device due to the lack of an efficient wireless powering approach. not exist. To compensate for poor signal conditions, sophisticated and bulky external readers are usually required. Sensor passivity can also limit detectable stimuli. The mid-field sources and implants described herein allow substantial amounts of power to be transferred from a palm-sized external module to a small implantable module located almost anywhere in the body. This enables many new sensing applications for continuous monitoring such as, for example, post-operative oxygen sensing in the heart and brain.

As another example, the systems and methods described herein may be used in wireless endoscopy. Current capsule endoscopies have limited battery life and sometimes provide inadequate small bowel examinations. This lifespan limitation can be addressed by the system described herein. Additionally, the implants described herein can be sized significantly smaller than current capsule endoscopy, allowing the patient to swallow multiple devices simultaneously. Each device can be oriented in a different direction within the intestine and thus can capture images from different angles at the same location, resulting in an improved field of view and improved diagnosis. Finally, the likelihood of retention may be reduced and the need for surgical or endoscopic retrieval may be avoided.

The systems and methods described herein can also be used for implanted drug delivery. Current implantable drug delivery systems are large and generally cannot be placed close enough to the site where the drug is needed. The implants described herein can further comprise one or more drug reservoirs. Implants can be injected or delivered to a target tissue area (eg, a tumor) via a catheter. A drug reservoir may be actuated to release drug by a midfield source. In some variations, this actuation may be controlled by the patient or clinician via a user interface, as described herein.

Additionally, the systems and methods described herein can be used in laboratory experiments using laboratory animals such as rodents (eg, mice, rats, etc.). The small size of the implant allows for monitoring capabilities that were not previously available or could not be readily implemented. For example, the implants described herein can be used to monitor or sense parameters and/or provide stimulation. Implants can be implanted on or near the brain of rodents, for example, to monitor electrical signals. The implant may be wirelessly powered by the midfield source described above and configured to communicate information back to the external module.

EXAMPLES The system described herein was used in two simulations of power transfer to an implant using a porcine tissue volume: the first simulated placement in the left ventricle of the heart, and the cortical region of the brain. was used in a second simulation arrangement in The source and implant were placed at least 5 cm apart. Figures 26A and 26B show magnetic resonance imaging (MRI) reconstructions of the implant location within the porcine tissue volume. FIG. 26A shows an MRI reconstruction of a configuration for transmitting power across the porcine chest wall to an implant placed on the heart surface. MRI images were acquired using a T2-weighted spin-echo pulse sequence and images were reconstructed using the OsiriX software package. The center of the source (white dot) and the coil of the implant (grey dot) were 5 cm apart (1 cm void, 4 cm inhomogeneous tissue). The fields in Figure 26A were calculated using a commercially available electromagnetic field simulator. A patterned metal plate was placed on 5 of the tissue multilayers (1 cm void, 4 mm skin, 8 mm fat, 8 mm muscle, 16 mm bone, 144 mm heart) and the fields were computed by a time-domain solver.

FIG. 26B shows an MRI reconstruction of the configuration for power transmission to an implant placed in the hypodermal layer region of the pig brain. MRI images were acquired using T2-weighted fast spin echo and images were reconstructed using the OsiriX software package. In the configuration shown there, the source-implant separation was 5.5 cm. When 500 mW (approximately the output power of a mobile phone) was coupled into the tissue, the power delivered to the coil of the implant was measured to be 195 μW for implants placed on the heart surface and 200 μW for implants placed on the hypodermal layer. . The received power remained substantially (˜10 μW) even when the operating depth (ie, the distance between the source and the implant) was increased to 10 cm.

These levels are far greater than the requirements of advanced integrated circuits. To illustrate the range of applications available with the performance characteristics reported in the main papers, Table 1 below shows the power requirements of state-of-the-art integrated circuits. The table is not exhaustive, but represents existing solid-state circuit capabilities in the microwatt power domain. Most of these devices are currently powered using either wire tethers or large (>2 cm) near-field coils.

Table 1: Selected Integrated Electronics Manufacturing Process and Power Consumption

By comparison, a cardiac pacemaker consumes approximately 8 μW. Computational studies show that performance is not affected by the microstructure and composition of the intermediate tissue, provided the field can be refocused.

Excess energy dissipated on tissue can pose potential safety concerns. A basic metric of RF exposure is the Specific Absorption Rate (SAR), defined as the integral of power loss over a reference volume of tissue. Limits exist on the SAR induced by electromagnetic field sources to prevent adverse health effects resulting from tissue heating. The system meets the IEEE C95.3-2005 standard if (i) the average whole body SAR is less than 0.4 W/kg and (ii) the maximum local SAR (averaged over 10 g of tissue) does not exceed 10 W/kg. Compliant. These limits are reduced by a factor of 5 for general public exposure (uncontrolled environment) such as mobile phones.

To assess exposure levels induced by power transfer, sources as described herein were operated on simulated tissue volumes defined by anthropomorphic glass fiber shells. The spatial distribution of the absorbed power was measured by scanning the robotic probe through a dosimetric liquid that mimics the body and head, as shown in Figure 27A. When 500 mW of focused power is coupled into tissue, the maximum specific absorption rate (SAR) is 0.89 W/kg in the body, averaged over 10 g of tissue, as shown in Figures 27B and 27C. , was found to be 1.17 W/kg in the head. These levels are well below the exposure thresholds of the controlled environment, as shown in Figure 27D. FIG. 27E shows that 2.2 mW and 1.7 mW are transmitted, respectively, for the configuration shown in FIGS. 26A and 26B if the power coupled to the tissue is allowed to meet the maximum permissible exposure level. It shows that it can be done. The low mean body absorption rate (less than 0.04 W/kg of adult body weight) and localized distribution suggest that power transfer is unlikely to have a significant effect on core body temperature. FIG. 27F shows that the power level is below the safe threshold.

Implantable mid-field receiver Conventional wireless power transfer techniques through tissue, where the transmitter and receiver are within the (in-air) wavelength of each other, dominate within the near-field of the transmitter and receiver structures. It relies on joins where different field types are the same. For example, an external transmitter loop can transmit a magnetic field that is inductively coupled to an embedded receiver loop via a magnetic field. In another example, an external dipole can be coupled with an electric field to an embedded dipole.

However, using a mid-field external transmitter, an electric field-based receiver (eg, dipole antenna) can be coupled to an electric field (eg, tangential H-field) based transmitter. With strong tangential H-field components, the magnetic field can propagate through the tissue medium. In mid-field transmitters the electric and magnetic fields are proportional in the induced propagating wave. A mid-field transmitter with a strong magnetic field component can be coupled to the electric field-based receiver.

A conventional receiver antenna coupled to a midfield transmitter includes a helical structure. In contrast to helical structures, which require three-dimensional fabrication techniques, dipoles can be readily fabricated, such as on planar surfaces. Also, dipoles can be more easily incorporated into injectable (eg, elongated) implants than helical structures.

FIG. 29 shows an implantable device 3000 implanted in tissue 3001. FIG. The illustrated implantable device 3000 includes a dipole antenna 3002 and receiver 3004 encapsulated within material 3006 and an outer implant casing 3010 . Implantable device 3000 further includes optional surface electrodes 3008 on outer implant casing 3010 . Surface electrodes 3008 are optional. For example, surface electrodes 3008 transmit received energy to tissue 3001 when apparatus 3000 is used as an ablation device. However, electrodes 3008 may not be necessary in applications where the receiver is used to help power electronics within an implanted device (eg, an implanted sensor).

Dipole antenna 3002 is made of a conductive material such as a metal, semiconductor, polymer, or other conductive material. The dipole antenna 3002 can include two straight thin conductors, as shown in FIG. 29, or folded dipoles, short dipoles, cage dipoles, bowtie dipoles, batwing dipoles, etc. may include other dipole antenna geometries. The use of geometries other than linear dipoles generally increases the width (eg, diameter) of device 3000 relative to the diameter of devices containing linear thin dipoles.

Receiver 3004 may be any receiver capable of receiving a signal from a midfield coupler. In one or more embodiments, the receiver may be an ultra-high frequency (UHF) receiver, such as capable of receiving signals transmitted at frequencies of approximately 2.45 GHz. The wavelength of such a signal in air is approximately 12.25 centimeters.

Material 3006 may be a high dielectric low loss material such as PREPERM®, polytetrafluoroethylene (PTFE) such as high dielectric PTFE, Eccostock® material, or RT/duroid®. Material 3006 may have a dielectric constant between that of midfield coupler substrate 1312 (see FIG. 13) and that of tissue 3001 in which material 3006 is embedded. Such a configuration makes the size of the receiver or receivers generally proportional to the wavelength of the signal incident on them, thus making the receivers larger than allowed for purely tissue-implanted receivers. be able to. Thus, a purely tissue-coupled receiver is compared to a receiver containing a dielectric with a dielectric constant between that of the midfield coupler substrate 1312 and that of the tissue 3001 in which the material 3006 is embedded. small. This can increase the efficiency of the power transmission link.

Consider an implantable receiver coupled to a small antenna encapsulated in a low dielectric material. The receiver is implanted in tissue with a large dielectric constant relative to a low dielectric constant encapsulant. A large power loss is realized between the high dielectric constant tissue and the low dielectric constant encapsulant. To reduce this loss, the material 3106 has a dielectric constant close to that of the surrounding tissue, e.g., to better match the perceived impedance of the tissue with the perceived impedance of the encapsulant. obtain. In general, using a high dielectric material as the encapsulant results in less change in the perceived impedance of the receiving circuit than using a low dielectric material as the encapsulant. That is, assume the impedance of the receiver in air is 'd1' and the impedance of the receiver in tissue is 'd2'. Also, the impedance difference |d2-d1|=Δ1 for a high impedance encapsulant and |d2-d1|=Δ2 for a low impedance encapsulant. In general, Δ2>Δ1. Smaller impedance changes allow for reduced dynamic range of adaptive impedance matching networks (eg, programmable inductors and/or programmable capacitors) at the receiver interface. Another advantage of using an encapsulant with a higher dielectric constant includes the receiver being immune to changes in the surrounding tissue dielectric properties that can occur due to the formation of scar tissue or adipose tissue.

Electrodes 3008 are optional, conductive elements that are electrically coupled to receiver 3004 . Electrodes 3008 transmit energy received at receptors 3004 (electric field energy) to tissue 3001 in contact with electrodes 3008 , eg, to ablate tissue 3001 .

Outer implant casing 3010 encloses dipole antenna 3002 , receiver 3004 and encapsulant 3006 . In one or more embodiments, the outer implant casing 3010 is made of polyurethane, silicone, ceramic, other urethane blends, Tecothane®, polyetheretherketone (PEEK), Pebax®, nylon, polycarbonate, It may consist of acrylonitrile butadiene styrene (ABS), thermoplastics, epoxies, combinations thereof, and the like.

Adjusting Transmitter Phase and/or Amplitude Previous solutions to help focus energy to implanted receivers include implanted power detectors, as described above. When using a time-domain multiplexed communication system between an external transmitter and an implanted receiver, the energy can be focused on the implanted receiver without using a power detector at the implant (e.g., energy can be more efficiently The phase and amplitude can be dynamically adjusted to help focus).

FIG. 30 shows a time domain multiplex communication system 3100 . The illustrated system 3100 includes an external midfield transceiver 3102 and an implantable transceiver 3104 . Transceiver 3102 includes a communicatively coupled mid-field antenna 3106 and transceiver 3104 includes a communicatively coupled electric field-based antenna 3108 . Antenna 3106 and antenna 3108 may be configured (eg, in length, width, shape, material, etc.) to transmit and receive signals at the same frequency. Transceiver 3104 may transmit data signals to transceiver 3102 via antenna 3108 and receive power and data signals transmitted by transceiver 3102 via antenna 3106 .

External midfield couplers (external transmitters) and implanted transceivers (including implanted antennas) can be used for both transmission and reception of RF signals. A T/R switch can be used to switch each RF port of the external transmitter from transmit (transmit data or power) mode to receive (receive data) mode (see FIG. 31). A T/R switch can be used to switch the implant between transmit (data transmit mode) and receive (power or data receive) modes (see FIG. 31).

The output of the receive terminal (on the external transmitter) of the T/R switch may be connected to one or more components that detect the phase and/or amplitude of the received signal from the implant. This phase and amplitude information can be used to program the phase of the transmitted signal to be substantially the same relative phase as the received signal. To help accomplish this, transceiver 3102 may include phase and amplitude matching network 3200, as shown in FIG. Network 3200 is for use with a midfield coupler that includes four ports, such as midfield coupler 602 of FIG. 6C. The illustrated network 3200 includes a midfield coupler 3202 electrically coupled to a plurality of switches 3204A, 3204B, 3204C, 3204D. Switches 3204A-3204D are each electrically coupled to phase and/or amplitude detectors 3206A, 3206B, 3206C, 3206D and variable gain amplifiers 3208A, 3208B, 3208C, 3208D, respectively. Amplifiers 3208A-3208D are electrically coupled to phase shifters 3210A, 3210B, 3210C, 3210D, respectively, which are power dividers that receive RF input signal 3214 transmitted through midfield coupler 3202. 3212.

Midfield coupler 3202 may be any midfield coupler described herein. Switches 3204A-3204D may be selector switches that select either the receive line (“R”) or the transmit line (“T”). The number of switches 3204A-3204D in network 3200 may equal the number of ports in midfield coupler 3202. FIG. In the example of network 3200, midfield coupler 3202 has four ports, but any number of one or more ports (and switches) can be used. In the example of a midfield coupler with a single port, power splitter 3212 is not required.

Phase and/or amplitude detectors 3206A-3206D detect the phase (Φ ₁ , Φ ₂ , Φ ₃ , Φ ₄ ) and power (P ₁ , P ₂ , P ₃ , P ₄ ). Phase and/or amplitude detectors 3206A-3206D may include one or more modules (hardware that may include electrical or electronic components configured to perform operations such as determining the phase or amplitude of a signal). module) and may include, for example, a phase detector module and/or an amplitude detector module. Detectors 3206A-3206D may include analog and/or digital components configured to determine the phase and/or amplitude of signals received at midfield detector 3202 .

Amplifiers 3208A-3208D take inputs (eg, M) from phase shifters 3210A-3210D (eg, _Pk phase-shifted by _Φ1 + _Φk , _Φ2 + _Φk , _Φ3 + _Φk , or _Φ4 + _Φk ). can receive. The amplifier output O is typically the power divider output M when the amplitude of the RF signal 3214 is 4*M in the example of FIG. 31, multiplied by the amplifier gain P _i *P _k . P _k can be dynamically set as the values of P ₁ , P ₂ , P ₃ , and/or P ₄ change. Φ _k is a constant. Phase shifters 3210A-3210D set the relative phases of the ports based on the phases from detectors 3206A-3206D.

Consider the situation where the transmit power required to be transmitted from midfield coupler 3202 is _Ptt . The RF signal supplied to power divider 3212 has a power of 4*M. The output of amplifier 3208A is generally M*P ₁ *P _k . Therefore, the power transmitted from the midfield coupler is M*( _P1 * _Pk + _P2 * _Pk + _P3 * _Pk + _P4 * _Pk )= _Ptt . Solving for P _k gives P _k =P _tt /(M*(P ₁ +P ₂ +P ₃ +P ₄ )).

The amplitude of the signal at each RF port can be transmitted at the same relative (scaled) amplitude as the signal received at the respective port of the midfield coupler coupled thereto. The gains of amplifiers 3208A-3208D can be further improved to account for losses between transmission and reception of signals from the midfield couplers. Consider the reception efficiency η= _Φir / _Ρtt . where P _ir is the power received at the embedded receiver. The efficiency (eg, maximum efficiency) when phase and amplitude tuning is performed can be estimated from the amplitude received from the implant's transmitter at the external midfield coupler. This estimate is
where P _it is the original power of the signal from the embedded transmitter. The power of the signal from the embedded transmitter can be communicated to the external transceiver 3102 as data from the embedded transceiver 3104 . The amplitude of the signals received at amplifiers 3108A-3108D may be scaled according to the determined efficiency to help ensure that the implant receives power to perform its programmed operation(s). Given the estimated link efficiency η and implant power (e.g., amplitude) requirements of P _ir ', P _k can be scaled as P _k =P _ir '/[η(P ₁ +P ₂ +P ₃ +P ₄ )].

Control signals, eg, phase and gain inputs, for phase shifters 3210A-3210D and amplifier 3208A, respectively, may be provided by processing circuitry not shown in FIG. Circuitry has been omitted so as not to overcomplicate or obscure the diagram shown in FIG. The same or different processing circuitry may be used to change switches 3204A-3204D from receive lines to transmit lines or vice versa. Again, this processing circuitry is not shown in FIG. 31 so as not to overcomplicate or obscure the illustration shown in FIG. See control unit 2112 in FIG. 21A for an example of such processing circuitry.

Customization of the dimension(s) of the midfield coupler Any body can be customized for structure (e.g., tissue, muscle density, fat content, cartilage, scar tissue, tendon composition, or other structural characteristics such as bone), contour, and/or the shape is different. Different mid-field coupler geometries have different properties that help the receiver in the body accept power transfer more efficiently and/or conform comfortably to the external surface of the body (e.g. skin like epidermal layer). can provide. A midfield coupler having a first shape may be more efficient in delivering power to the first body, but less efficient in delivering power to the second body.

FIG. 32 shows a midfield coupler applied to tissue (eg, human or other animal skin). Midfield coupler system 3300 as shown includes midfield coupler 3302 , electronic module 3304 , RF connectors 3306 A, 3306 B, and molded backing layer 3308 . System 3300 is shown attached to tissue 3310 .

The operations illustrated in FIG. 33 illustrate how a midfield coupler can be designed to accommodate a particular body shape, contour, and/or structure. This procedure is useful when high field strength is required to power more power-hungry electronics or to perform ablation on the target area, and when implants are deeper than near-field. For example, it allows the design of a focused midfield transmitter to efficiently power or energize the anatomy, such as when implanted in the midfield.

At operation 3402, an anatomical structure (eg, structure 3310) may be imaged, such as using a magnetic resonance imaging (MRI) device, computed tomography (CT) device, or other imaging device. The anatomy includes the area in which the implant will be placed. At operation 3404, the imaged structure is a simplified model of the geometry of materials with known dielectric properties (eg, tissue, bone, tendon, cartilage, scar tissue, organs, fluids and/or blood vessels, etc.). can be decomposed into At operation 3406, the current distribution at the target frequency (eg, 915 MHz, 2.45 GHz, or other microwave frequency) is determined. At operation 3408, the dimensions of the midfield coupler that can deliver energy at or near the determined current distribution (e.g., strip width/length, slot width/length, spacing between multiple slots, impedance matching). The values of one or more passive components (eg, programmable passive components such as capacitors or inductors) that may be used, or ports of a midfield source such as midfield coupler 3302, are determined. Not all current distributions are possible, so it may be necessary to choose a different implantation site or operate the midfield coupler at less than optimum efficiency. The current distribution can be determined by solving the current distribution equations discussed above.

At operation 3410, a customized midfield plate for the midfield coupler may be fabricated (eg, etched, plated, and/or printed). The plate can be made by using standard manufacturing techniques and/or materials, which can include, for example, FR4, polyimide, or other materials. At operation 3412, an electrical/electronic component can be electrically coupled to the midfield coupler. Components include one or more connectors, such as RF connectors 3306A-3306B, an electrical/electronic module 3304 (eg, one or more transistors, resistors, capacitors, transceivers (eg, radio and antenna transmit/receive)). , inductors, digital logic such as logic gates (eg, programmable logic gates), arithmetic logic units (ALUs), processors, and the like. Electrical/electronic module 3304 may include switches 3204A-3204D, detectors 3206A-3206D, amplifiers 3208A-3208D, phase shifters 3210A-3210D, and/or power divider 3212.

At operation 3414 , an insulating material such as material 3308 may be applied to midfield coupler 3302 . Insulating materials may include foams, polymers (eg, plastics), or silicones. A substance 3308 may be applied to the surface of midfield coupler 3302 to form an insulating layer between the midfield coupler and tissue 3310 . Material 3308 can be shaped, cut, or 3D printed to conform to the shape of the skin surface. Imaging the surface contour can be done using a camera, laser, or cast. Conforming the material to the tissue 3310 can enhance patient comfort while mounting the transmitter and minimizing slippage or displacement of the transmitter from the target anatomy. Minimizing slip can be important for non-adjustable midfield couplers where the focal region of the transmitter is fixed. Additional backing material may be added to create a soft interface between the skin and the midfield coupler unit.

Other Applications One or more of the systems, devices, and methods described herein may be used in the tibial nerve or any nerve branch of the tibial nerve (eg, but not limited to, the posterior tibial nerve), the sacral nerve. Faecal or urinary incontinence (e.g., excessive It can be used to help treat active bladder). One or more of the systems, devices, and methods described herein can be used to help treat urinary incontinence, such as overactive bladder. Urinary incontinence is caused by stimulating one or more of the muscles of the pelvic floor, the nerves that innervate the muscles of the pelvic floor, the internal urinary sphincter, the external urethral sphincter, and the pudendal nerve or nerve branches of the pudendal nerve. It can be treated using wireless transmission.

One or more of the systems, devices, and methods described herein may include nerves or branches of the hypoglossal nerve, the base of the tongue (muscle), the phrenic nerve (singular), the intercostal nerve (singular), ), accessory nerve(s), and cervical nerves C3-C6 to help treat sleep apnea and/or snoring. The treatment of sleep apnea and/or snoring is to use a mid-field coupler to energize the implant to detect decreased, impaired, or stopped breathing (e.g., by measuring oxygen saturation). can include

One or more of the systems, devices and methods described herein may be used, for example, in one or more of the Bartholin's gland(s), Skene's gland(s), and the inner wall of the vagina. Multiple stimulation may be used to help treat vaginal dryness. One or more of the systems, devices, and methods described herein may be used, for example, in the occipital nerve, the supraorbital nerve, the C2 cervical nerve or branches thereof, and the frontal nerve or branches thereof. It can be used to help treat migraines by stimulating one or more. One or more of the systems, devices, and methods described herein may, for example, stimulate the stellate ganglion and one or more of the sympathetic chains C4-C7 to stimulate the heart. It may be used to help treat post-traumatic stress disorder, hot flashes, and/or complex regional pain syndrome.

One or more of the systems, devices, and methods described herein stimulate, for example, the pterygopalatine ganglion block, the trigeminal nerve, or one or more of the nerve branches of the trigeminal nerve. It can be used to help treat trigeminal neuralgia by: One or more of the systems, devices, and methods described herein can be used, for example, in tissue of the parotid, submandibular, sublingual, buccal, labial and/or lingual mucosa. submucosa of the oral mucosa, soft palate, lateral part of the hard palate, and/or floor of the mouth, and/or intermuscular fibers of the tongue, von Ebner glands, glossopharyngeal nerve (CN IX) (CN IX including ear ganglion) facial nerve (CN VII) (including nerve branches of CN VII such as the submandibular ganglion and nerve branches of T1-T3 such as the superior cervical ganglion) can be used to help treat xerostomia (e.g., caused by side effects of medications, chemotherapy or radiotherapy cancer treatments, Sjögren's disease, or other causes of xerostomia) by stimulating the

One or more of the systems, devices, and methods described herein sense, for example, electrical output from a proximal portion of a severed nerve and output a signal to a distal portion of the severed nerve. of the severed nerve by delivering electrical input and/or by sensing electrical output from the distal portion of the severed nerve and delivering electrical input to the proximal portion of the severed nerve. It can be used to aid in therapy. One or more of the systems, devices, and methods described herein can be used, for example, in one or more muscles of a patient with cerebral palsy, or one or more muscles in which one or more muscles are affected. Or it can be used to help treat cerebral palsy by stimulating multiple innervations. One or more of the systems, devices, and methods described herein may be used, for example, in the pelvic visceral nerves (S2-S4) or any nerve branch thereof, the pudendal nerve, the cavernous nerve(s), ), may be used to help treat erectile dysfunction by stimulating one or more of the hypogastric plexus.

One or more of the systems, devices, and methods described herein can be used to help treat menstrual cramps, for example, by stimulating one or more of the uterus and vagina . One or more of the systems, devices, and methods described herein sense one or more of PH and blood flow, e.g., to aid in contraception, pregnancy, bleeding, or pain. or by delivering electrical current or drugs. One or more of the systems, devices, and methods described herein may, for example, stimulate the female genitalia (including external and internal including the female clitoris or other sensory active parts). , or to induce sexual arousal in humans by stimulating the male genitalia. One or more of the systems, devices, and methods described herein, for example, by stimulating one or more of the carotid sinus, the vagus nerve, or nerve branches of the vagus nerve, It can be used to help treat high blood pressure. By stimulating one or more of the trigeminal nerve or its branches, the anterior ethmoid nerve, and the vagus nerve, one or more of the systems, devices, and methods described herein provide , may be used to help treat paroxysmal supraventricular tachycardia.

One or more of the systems, devices, and methods described herein may be used, for example, by sensing activity in the vocal cords and the contralateral vocal cords, or by detecting the nerve innervating the vocal cords, the left recurrent laryngeal nerve. and/or stimulation of the right recurrent laryngeal nerve, and the vagus nerve to help treat vocal cord dysfunction by stimulating just one or more of the vocal cords. One or more of the systems, devices, and methods described herein stimulate tissue to perform one or more of microcirculatory stimulation and protein synthesis, e.g., to heal wounds; It can be used to aid in tissue repair by restoring connective tissue and/or dermal tissue integrity. One or more of the systems, devices, and methods described herein, for example, stimulate the vagus nerve or its nerve branches, block the release of norepinephrine and/or acetylcholine, and/or and/or to help with asthma or chronic obstructive pulmonary disease by one or more of blocking the receptor for acetylcholine. One or more of the systems, devices, and methods described herein can, for example, stimulate one or more nerves near or within a tumor to modulate, e.g., epinephrine By reducing sympathetic innervation such as /NE release and/or parasympathetic innervation such as Ach, it can be used to help treat cancer. One or more of the systems, devices, and methods described herein power sensors within the human body that detect diabetic parameters, such as blood glucose or ketone levels, for example. Such sensor data can be used to help treat diabetes by adjusting the delivery of exogenous insulin from an insulin pump. One or more of the systems, devices, and methods described herein power sensors inside the human body that detect diabetic parameters, such as, for example, blood glucose or ketone levels; It can be used to help treat diabetes by stimulating the release of insulin from pancreatic islet β-cells using field couplers.

ADDITIONAL EXAMPLES Example 1 describes a subject matter (e.g., an apparatus, method, means for performing an action, or a machine including instructions that when executed by a machine can be configured to cause the machine to perform the action). a readable memory), for example, a first transceiver for transmitting and receiving microwave signals at a first frequency, the signal from the first transceiver being transmitted parallel to the surface of the midfield coupler; a first transceiver including a midfield coupler that converts the signal into a signal having a non-negligible magnetic field component and focuses the converted signal to a location within the tissue that is within the wavelength of the microwave signal when measured in air; , an at least partially implantable biocompatible device comprising an electric field-based antenna that receives signals from the midfield coupler, and a second transceiver that transmits signals at approximately the same frequency as the first transceiver. obtain or use.

Example 2 may include or use the subject matter of Example 1, or may optionally be combined with the subject matter of Example 1 to include or use the subject matter of Example 1. In this case the electric field based antenna is a dipole antenna.

Example 3 may include or use the subject matter of at least one of Examples 1-2, or optionally includes or use the subject matter of at least one of Examples 1-2. can be combined with the subject matter of at least one of Examples 1-2 to do so. In this case, the first transceiver comprises a phase matching network comprising a phase detector and a phase shifter electrically coupled to the midfield coupler, the phase detector determining the phase of the signal received from the second transceiver. and a phase shifter adjusts the phase of the signal provided to the midfield coupler based on the determined phase of the signal received from the second transceiver.

Example 4 may include or use the subject matter of Example 3, or may optionally be combined with the subject matter of Example 3 to include or use the subject matter of Example 3. In this case, the phase shifter adjusts the phase of the signal according to the determined phase of the signal received from the second transceiver.

Example 5 may include or use the subject matter of Example 3, or may optionally be combined with the subject matter of Example 3 to include or use the subject matter of Example 3. In this case, the phase shifter adjusts the phase of the signal to match the phase of the signal received from the second transceiver.

Example 6 may include or use the subject matter of at least one of Examples 1-5, or optionally includes or use the subject matter of at least one of Examples 1-5. can be combined with the subject matter of at least one of Examples 1-5 to do so. In this case, the first transceiver comprises an amplitude matching network comprising an amplitude detector electrically coupled to the midfield coupler and a variable gain amplifier, the amplitude detector for measuring the signal received from the second transceiver. Amplitude is determined and the variable gain amplifier adjusts the amplitude of the signal provided to the midfield coupler based on the amplitude of the signal received from the second transceiver.

Example 7 may include or use the subject matter of Example 6, or may optionally be combined with the subject matter of Example 6 to include or use the subject matter of Example 6. in this case the midfield coupler includes two or more ports, the amplitude detector is one of the two or more amplitude detectors electrically coupled to respective ports of the midfield coupler; The first transceiver receives a radio frequency (RF) signal and further comprises a power splitter for splitting the RF signal into two or more signals, one signal for each port of the midfield coupler; The gain amplifier is one of the plurality of variable gain amplifiers, each variable gain amplifier electrically coupled between a respective port of the midfield coupler and the power splitter, each amplifier comprising: receive one of the two or more signals from the power splitter and amplify the signal by a gain, the gain being the amplitude determined by the same individually coupled amplitude detector of the midfield coupler; determined based on

Example 8 may include or use the subject matter of Example 7, or may optionally be combined with the subject matter of Example 7 to include or use the subject matter of Example 7. In this case, the gain of each amplifier of the plurality of amplifiers is the amplitude determined by the amplitude detector multiplied by a constant amount.

Example 9 may include or use the subject matter of Example 8, or may optionally be combined with the subject matter of Example 8 to include or use the subject matter of Example 8. In this case the quantities are (27) where P _tt is a particular amplitude and P _i is the number of amplitudes determined in the amplitude detector for each of the i-ports of the midfield coupler. is the amplitude.

Example 10 may include or use the subject matter of Example 9, or may optionally be combined with the subject matter of Example 9 to include or use the subject matter of Example 9. In this case the quantity Pk is further divided by the efficiency index η, where P _it is the amplitude of the signal transmitted from the second transceiver.

Example 11 can include or use the subject matter of at least one of Examples 1-10, or optionally includes or use the subject matter of at least one of Examples 1-10 can be combined with the subject matter of at least one of Examples 1-10 to do so. In this case, the antenna is encapsulated in a dielectric material having a dielectric constant between that of animal tissue and that of the substrate of the midfield coupler on which the midfield plate of the midfield coupler is located.

Example 12 may include subject matter (e.g., an apparatus, method, means for performing operations, or a machine-readable memory containing instructions that, when executed by a machine, may configure a machine to perform actions), or, for example, a radio that transmits and receives microwave signals, and a midfield coupler electrically coupled to the radio, such that the signal from the radio is not negligible parallel to the surface of the midfield coupler. a midfield coupler for converting a signal having a magnetic field component and focusing the signal to a location within the tissue that is within the wavelength of the microwave signal as measured in air; an amplitude detector for determining the amplitude of a signal received at the midfield coupler; and a variable gain amplifier electrically coupled between the radio and the midfield coupler, and a variable gain amplifier that amplifies the transmitted signal from the to an amplitude determined by the amplitude detector.

Example 13 may include or use the subject matter of Example 12, or may optionally be combined with the subject matter of Example 12 to include or use a phase matching network comprising a phase detector and a phase shifter. . A phase shifter and a phase detector are electrically coupled to the midfield coupler, the phase detector determines the phase of the signal received at the midfield coupler, and the phase shifter determines the midfield signal based on the determined phase. Adjust the phase of the signal provided to the field coupler.

Example 14 may include or use the subject matter of Example 13, or may optionally be combined with the subject matter of Example 13 to include or use the subject matter of Example 13. In this case, the phase shifter adjusts the phase of the signal by the determined phase.

Example 15 may include or use the subject matter of at least one of Examples 12-14, or optionally include or use the subject matter of at least one of Examples 12-14 can be combined with the subject matter of at least one of Examples 12-14 to do so. in this case the midfield coupler includes two or more ports, the amplitude detector is one of the two or more amplitude detectors electrically coupled to respective ports of the midfield coupler; The first transceiver receives a radio frequency (RF) signal and further comprises a power splitter for splitting the RF signal into two or more signals, one signal for each port of the midfield coupler; The gain amplifier is one of the plurality of variable gain amplifiers, each variable gain amplifier electrically coupled between a respective port of the midfield coupler and the power splitter, each amplifier comprising: Receives one of the two or more signals from the power splitter and amplifies the signal by a gain, the gain being determined based on the amplitude determined by the amplitude detector.

Example 16 may include or use the subject matter of Example 15, or may optionally be combined with the subject matter of Example 15 to include or use the subject matter of Example 15. In this case, the gain of each amplifier of the plurality of amplifiers is the amplitude determined by the amplitude detector multiplied by a constant amount.

Example 17 may include or use the subject matter of Example 16, or may optionally be combined with the subject matter of Example 16 to include or use the subject matter of Example 16. 29, where P _tt is a specific amplitude and P _i are the multiple amplitudes determined in the amplitude detector for each of the i-ports of the midfield coupler. , and where P _it is the amplitude of the signal sent to the midfield coupler.

Example 18 may include subject matter such as an apparatus, method, means for performing operations, or a machine-readable memory containing instructions that, when executed by a machine, may be configured to cause the machine to perform actions. or, for example, an at least partially implantable biocompatible device comprising an outer casing; a radio for transmitting and receiving microwave signals covered by the outer casing; An electric field-based antenna coupled and encased in an outer casing, and an encapsulant within the outer casing, surrounding the radio and antenna, between the permittivity of animal tissue and the permittivity of the substrate of the midfield coupler. An encapsulant that includes a dielectric constant may be included or used.

Example 19 may include or use the subject matter of Example 18, or may optionally be combined with the subject matter of Example 18 to include or use the subject matter of Example 18. In this case the antenna is a dipole antenna.

Example 20 may include or use the subject matter of at least one of Examples 18-19, or optionally one or more exposed on the casing and electrically coupled to the radio can be combined with the subject matter of at least one of Examples 18-19 to include or use electrodes of

Claims (20)

A first transceiver for transmitting and receiving microwave signals at a first frequency, the first transceiver including a midfield coupler for transmitting signals from the first transceiver parallel to a surface of the midfield coupler. a first transceiver that converts a signal having a magnetic field component that cannot be detected and focuses the converted signal to a location within tissue that is within the wavelength of the microwave signal when measured in air;
An at least partially implantable biocompatible device comprising an electric field-based antenna for receiving said signal from said midfield coupler and comprising a second transceiver for transmitting a signal at approximately the same frequency as said first transceiver. A system comprising and . 3. The system of claim 1, wherein the electric field-based antenna is a dipole antenna. The first transceiver comprises a phase matching network comprising a phase detector and a phase shifter electrically coupled to the midfield coupler, the phase detector detecting the phase of the signal received from the second transceiver. and the phase shifter adjusts the phase of the signal supplied to the midfield coupler based on the determined phase of the signal received from the second transceiver. system. 4. The system of claim 3, wherein said phase shifter adjusts said phase of said signal by said determined phase of said signal received from said second transceiver. 4. The system of claim 3, wherein said phase shifter adjusts said phase of said signal to match said phase of said signal received from said second transceiver. The first transceiver comprises an amplitude matching network comprising an amplitude detector and a variable gain amplifier electrically coupled to the midfield coupler, the amplitude detector being adapted to detect signals received from the second transceiver. 2. The method of claim 1, wherein an amplitude is determined and the variable gain amplifier adjusts the amplitude of the signal supplied to the midfield coupler based on the amplitude of the signal received from the second transceiver. system. the midfield coupler includes two or more ports;
wherein said amplitude detector is one of two or more amplitude detectors, each of said two or more amplitude detectors electrically coupled to respective ports of said midfield coupler;
the first transceiver receives a radio frequency (RF) signal, splits and separates the RF signal into two or more signals, one signal corresponding to each port of the midfield coupler; further equipped with a distributor,
The variable gain amplifier is one of a plurality of variable gain amplifiers, each variable gain amplifier electrically coupled between respective ports of the midfield coupler and the power divider. , each amplifier receives one of said two or more signals from said power splitter and amplifies said signal by a gain, said gain being equal to each of said midfield couplers; 7. The system of claim 6, determined based on amplitudes determined by the coupled amplitude detectors. 8. The system of claim 7, wherein the gain of each amplifier of the plurality of amplifiers is the amplitude determined by the amplitude detector multiplied by an amount. Said amount is
where P _tt is a particular amplitude and P _i is an amplitude among multiple amplitudes determined in the amplitude detector for each of the i-ports of the midfield coupler. Item 9. The system according to Item 8. Said quantity Pk is further divided by the efficiency index η, where:
and P _it is the amplitude of the signal transmitted from the second transceiver. 2. The antenna of claim 1, wherein the antenna is encapsulated in a dielectric material having a dielectric constant between the dielectric constant of animal tissue and the dielectric constant of the substrate of the midfield coupler on which the midfield plate of the midfield coupler is disposed. System as described. a radio for transmitting and receiving microwave signals;
A midfield coupler electrically coupled to the radio for converting a signal from the radio into a signal having a non-negligible magnetic field component parallel to the surface of the midfield coupler and measured in air. a mid-field coupler that focuses the signal to a location within tissue that is within the wavelength of the microwave signal when
an amplitude detector electrically coupled to the midfield coupler for determining the amplitude of a signal received at the midfield coupler;
A variable gain amplifier electrically coupled between the radio and the midfield coupler for amplifying a transmitted signal from the radio to the amplitude determined by the amplitude detector. An apparatus comprising an amplifier and further comprising a phase matching network comprising a phase detector and a phase shifter;
The phase shifter and the phase detector are electrically coupled to the midfield coupler, the phase detector determining the phase of a signal received at the midfield coupler, the phase shifter 13. The apparatus of claim 12, adjusting the phase of the signal supplied to the midfield coupler based on the determined phase. 14. The apparatus of claim 13, wherein said phase shifter adjusts said phase of said signal by said determined phase. the midfield coupler includes two or more ports;
wherein said amplitude detector is one of two or more amplitude detectors, each of said two or more amplitude detectors electrically coupled to respective ports of said midfield coupler;
The first transceiver receives a radio frequency (RF) signal, splits and separates the RF signal into two or more signals, and power distributes one signal to each port of the midfield coupler. with more equipment,
The variable gain amplifier is one of a plurality of variable gain amplifiers, each variable gain amplifier electrically coupled between respective ports of the midfield coupler and the power divider. , each amplifier receives one of the two or more signals from the power splitter and amplifies the signal by a gain, the gain determined by the amplitude detector 13. The apparatus of claim 12, determined based on amplitude. 16. The apparatus of claim 15, wherein the gain of each amplifier of the plurality of amplifiers is the amplitude determined by the amplitude detector multiplied by an amount. Said amount is
where P _tt is a particular amplitude, P _i is an amplitude of a plurality of amplitudes determined in the amplitude detector for each of the i-ports of the midfield coupler, and ,
and P _it is the amplitude of the signal transmitted to the midfield coupler. An at least partially implantable biocompatible device comprising:
an outer casing;
a radio for transmitting and receiving microwave signals covered by the outer casing;
an electric field-based antenna electrically coupled to the radio and covered by the outer casing;
an encapsulant within the outer casing surrounding the radio and the antenna and having a dielectric constant between that of animal tissue and that of a midfield coupler substrate. 19. The apparatus of claim 18, wherein said antenna is a dipole antenna. 19. The apparatus of claim 18, further comprising one or more electrodes exposed on said outer casing and electrically coupled to said radio.