Patent Publication Number: US-10307607-B2

Title: Focused magnetic stimulation for modulation of nerve circuits

Description:
TECHNICAL FIELD 
     This disclosure relates generally to devices, systems, and methods for modulation of nerve circuits by focused magnetic stimulation. 
     BACKGROUND 
     Neuromodulation is an evolving therapy that can involve various types of electromagnetic stimuli including the application of a strong magnetic field or a small electric current to nerve structures. 
     SUMMARY 
     Some embodiments are directed to a neuromodulation device that includes electrically conductive coils arranged in an array and circuitry coupled to energize the coils in the array using current pulses that generate an electromagnetic field. The circuitry is configured to control one or more parameters of the current pulses, including at least amplitude and phase of the current pulses, such that the electromagnetic field undergoes constructive and destructive interference that focuses and/or steers a magnetic flux density within a region of interest of a patient. 
     Some embodiments involve a neuromodulation system. The system includes a neuromodulation device comprising electrically conductive coils arranged in an array and circuitry coupled to energize the coils in the array with current pulses that generate an electromagnetic field. The circuitry is configured to control one or more parameters of the current pulses, including at least amplitude and phase of the current pulses, such that the electromagnetic field undergoes constructive and destructive interference that focuses and/or steers a magnetic flux density within a region of interest of a patient. Communications circuitry is configured to wirelessly transfer communication signals between the neuromodulation device and an external device. A patient information device is communicatively coupled to the neuromodulation device and is configured to monitor one or more biological signals of the patient and to transfer information about the biological signals to the neuromodulation device via the communication signals. 
     A neuromodulation method includes energizing coils in an array of coils using current pulses. The current pulses generate an a electromagnetic field that provides a magnetic flux density at a region of interest of a patient. One or more parameters of the current pulses, including at least amplitude and phase, are controlled such that the electromagnetic field undergoes constructive and destructive interference that focuses and/or steers the magnetic flux density to the region of interest. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1A  is a block diagram of a neuromodulation device in accordance with some embodiments; 
         FIG. 1B  illustrates a portion of the circuitry and coil array in accordance with some embodiments. 
         FIG. 1C  illustrates the driver circuitry for each coil in accordance with various embodiments; 
         FIG. 1D  illustrates the magnetic field created by a single loop coil; 
         FIG. 1E  illustrates the magnetic field created by a multi-loop coil; 
         FIG. 1F  illustrates manipulation the field strength of the electric field at a selected location by approximately adding the linear vectors of the individual fields in accordance with some embodiments; 
         FIG. 2  illustrates an embodiment of the neuromodulation device wherein the device is disposed on a flexible substrate configured as an implantable nerve cuff in accordance with some embodiments; 
         FIG. 3  illustrates a two dimensional flat planar coil that may be used in the coil array the neuromodulation device in accordance with some embodiments; 
         FIG. 4  shows a three dimensional coil that may be used in the coil array the neuromodulation device in accordance with some embodiments; 
         FIG. 5A  shows a measured stress versus sputter pressure plot for MoCr; 
         FIGS. 5B and 5C  illustrate a process for forming a stress engineered coil structure in accordance with some embodiments; 
         FIGS. 6A through 6C  illustrate a process for forming a stress engineered coil structure in accordance with some embodiments; 
         FIG. 6D  shows interlocking tips of the coil loops in accordance with some embodiments; 
         FIGS. 7A through 7C  illustrate focused magnetic stimulation that is enabled by the neuromodulation devices and systems as described in embodiments herein; 
         FIG. 8  is a plot of focal depth vs. spot size that is achievable using the neuromodulation devices according to embodiments described herein; and 
         FIGS. 9A and 9B  show a patient-internal neuromodulation device used in conjunction with a patient-external patient information device in accordance with some embodiments. 
     
    
    
     The figures are not necessarily to scale. Like numbers used in the figures refer to like components. However, it will be understood that the use of a number to refer to a component in a given figure is not intended to limit the component in another figure labeled with the same number. 
     DETAILED DESCRIPTION 
     Up to 23% of patients with nerve stimulation devices experience surgery-related and/or other complications, hardware malfunctions, and/or adverse side effects. Many of these adverse outcomes are the result of one or more of the relatively large size of the nerve stimulation device, the invasive nature of surgical implantation, the relatively large, unfocused spread of electrical currents into off-target regions because of the inability of electric stimulation to reliably activate specific sections of the nerve, and the lack of a sensor-driven smart algorithms to provide feedback control to optimize stimulation. Current neuromodulation technologies stimulate large volumes including unwanted regions, and may not penetrate sufficiently below the skin surface. These limitations hinder the applications and effectiveness of classic neuromodulation technology, not only for brain stimulation, but also for modulating peripheral nerve circuits. 
     Approaches discussed herein are directed to high-precision spatial targeting of nerve circuits by shaping magnetic fields. The neuromodulation devices disclosed herein provide minimally-invasive and/or feedback-controlled neural modulation for regulating brain stimulation as well as regulating peripheral nerve circuits such as the vagus nerve. The ability to selectively stimulate nerve fascicles enables treatment of a wide-range of peripheral and central nervous system disorders with targeted therapies. The focused magnetic stimulation (FMS) neuromodulation approaches disclosed herein are underpinned by metamaterial coils as discussed below. These micro-engineered metamaterial structures allow for far greater control of electromagnetic fields over conventional transducer technologies. Driven by smart current distribution algorithms, FMS can non-invasively target small bundles of nerve fibers, as well as provide tailored stimulus patterns. The use of an array of metamaterial coils combined with a current distribution algorithm enables more localized stimulations, deeper penetration, enhanced depth control, and complex stimulation patterns with the ability to target specific nerve fascicles. 
     Turning now to  FIG. 1  there is shown a neuromodulation device  100  configured to provide neuromodulation stimulation n. The neuromodulation device  100  may be useful for transcranial nerve stimulation as well as stimulation of the peripheral nerves, such as the vagus nerve, for example. The device  100  includes electrically conductive coils  110   a  arranged in an array  110  and circuitry  120 . 
     The circuitry  120  includes multiplexer circuitry  120   a  configured to allow access to individual coils of the array  110 ; a controller  120   b  configured to control the parameters of the current pulses provided to energize the coils  110   a;  power management circuitry  120   c  configured to provide power for the current pulses; and driver circuitry  120   f  configured to energize the coils  110   a  in the array  110 . In some embodiments, the circuitry  120  may include a battery and/or energy harvesting circuitry  120   d  that supplies energy to the power management circuitry  120   c  and communication circuitry configured to communicate with an external device  150 . 
     The controller  120   b  is configured to control one or more parameters of the current pulses, including at least amplitude and phase of the current pulses, such that the electromagnetic fields produced by the coils  110   a  in the array  110  undergo constructive and destructive interference that focuses and/or steers a magnetic flux density within a region of interest  105  of a patient  101 . In some embodiments, a neuromodulation device  100  may be an external therapy system that is placed on or above the skin  101   a  of the patient  101  as illustrated in  FIG. 1A . In some embodiments, the neuromodulation device  100  may be at least partially implantable. For example, the coil array  110  and/or circuitry  120  may be implanted subcutaneously. 
     In some implementations, in addition the control of the amplitude and phase of the current pulses, the controller  120   b  may be configured to additionally additional parameters of the current pulses such as the duty cycle and/or frequency of the current pulses. Control of the current pulse parameters is used to focus and/or steer the magnetic flux density within the region of interest  105 . 
     The coils  110   a  of the array  110  may have a diameter in a range of greater than or equal to about 100 μm to about 500 μm, or in a range of greater than or equal to about 10 μm to about 100 μm, for example. In some embodiments, the coils  110   a  are 2D planar coils and in some embodiments, the coils are 3D metamaterial coils made of one or more stressed elastic members as disclosed in commonly owned U.S. Pat. No. 6,646,533 which is incorporated by reference herein. The resolution of the stimulation head comprising array  110  may be about three times the diameter of one of the coils  110   a  in the array  110 . In one example implementation, the neuromodulation device may comprise a 10×10 array of coils, each coil having a diameter of 150 μm, a coil pitch (center to center distance between coils) of 150 μm, a maximum stimulation area size of 3 mm 2 , and current injection in each coil of less than 100 mA. 
     In some embodiments, the magnetic flux density within the region of interest  105  is greater than about 0.1 Tesla, the electric field strength within the area of interest  105  may be about Ex=dV/dx&gt;100 V/m, an electric field gradient within the area of interest  105  may be about dEx/dx&gt;500 V/m 2  and/or a maximum electric current pulse amplitude in each coil may be less than about 500 mA or even less than about 100 mA. In some implementations, the magnetic flux density, electric field strength and/or electric field gradient produced by the array  110  is sufficient to activate one or more neurons within the region of interest  105  to provide neuromodulation therapy. For example, the magnetic flux density, electric field strength, and/or electric field gradient produced by the array  110  within the region of interest  105  may be sufficient to activate a nerve fascicle at a specified depth within a nerve bundle while not activating other nerve fascicles of the nerve bundle. In some scenarios, the neuromodulation therapy may involve using the array  110  to provide a magnetic flux density, electric field strength, and/or electric field gradient at a sub threshold level that is below the activation threshold of the nerve fibers in the region of interest. 
     The region of interest  105  is located at a specified depth within the patient  101  and the magnetic flux density, electric field strength, and/or electric field gradient produced by the neuromodulation device  100  in a region  106  between the region of interest  105  and the array  110  is less than the magnetic flux density, electric field strength, and/or electric field gradient in the region of interest  105 . 
     Optionally, the neuromodulation device  100  includes a substrate  130 , wherein the array  110  of coils  110   a  and the circuitry  120  are disposed on the substrate  130 . The substrate  130  can be flexible. For example, in various embodiments, the substrate  130  may comprise an implantable nerve cuff or a dermal patch. In some implementations, the neuromodulation device can be printed on a flexible substrate. 
     In some embodiments, the neuromodulation device  100  includes power supply circuitry that optionally comprises a battery. In some embodiments, the power management circuitry  120   c  obtains power from a power supply  120   d  such as a battery or energy harvesting circuit. When present, the energy harvesting circuit is configured to harvest power from a radio frequency (RF) signal generated by an additional device  150  which may be a patient-external device. The power supply  120   d  provides power to the power management circuitry  120   c  which uses the harvested power to provide the current pulses to the coils  110   a.    
     In some embodiments, the neuromodulation device includes communications circuitry  120   e  configured to wirelessly transfer communication signals between the neuromodulation device  100  and an additional device  150 . The device  150  may be configured to obtain biological information from the patient wherein the biological information is used to develop feedback information for the FMS. In some scenarios, the communications signals passed between the device  150  and the communications circuitry  120   e  include the feedback information and the controller  120   b  uses the feedback information to control the current pulse parameters. 
     In some embodiments, the controller  120   b  includes a memory that stores one or more tables of current pulse parameter values for each coil in the array, each table corresponding to a particular profile of biological information that is consistent with the current physiological state of the patient. The current physiological state of the patient may be provided to the controller  120   b  by the external device  150  via the communications circuitry  120   e . The controller  120   b  accesses the memory to retrieve the current pulse parameter values to be used for the stimulation, wherein the current pulse parameter correspond to the patient&#39;s physiological state. In some implementations, the current pulse parameter values utilized by the neuromodulation device are dynamically changeable in response to a change in the biological information obtained from the patient via device  150 . The circuitry  120  may be implemented using a silicon based application specific integrated circuit (ASIC) and/or a thin-film-transistor (TFT) circuitry backplane. TFT implementation is particularly useful for flexible substrates. 
     Each coil  110   a  in the array  110  may be individually addressable, e.g., using multiplexers  120   a  and the controller  230   b , to implement an addressable array. Each coil  110   a  is coupled to driving circuitry  120   f . In some implementations, the driving circuitry  120   f  may support bipolar currents by incorporating a pair of complementary transistors as shown in  FIG. 1C . Driving circuitry  120   f  allows programming of a distinct current through each coil  110   a  before the entire array  110  is activated. 
       FIG. 1B  illustrates a portion of the circuitry  120  and coil array  110  in accordance with some embodiments.  FIG. 1C  illustrates the driver circuitry  120   f  for each coil  110   a . As illustrated in  FIG. 1B , circuitry  120  includes a controller  120   b  and row and column multiplexers  120   a  configured to access the array coils  110   a . The circuitry  120  further includes power management circuitry  120   c  that operates in conjunction with the controller  120   b  and multiplexers  120   a  to implement the current pulse injection algorithm. The current pulse injection algorithm provides the current distribution in the coils  110   a  to enable high-precision targeting and/or to tailor stimulus patterns. For example, the current pulse injection algorithm may implement a phased array stimulation wherein coils  110   a  are selectively energized by current pulses that are in phase and/or out of phase to provide constructive and/or destructive interference between the electric fields generated by at least some of the coils  110   a  of the array  110 . The constructive and/or destructive interference in the electric fields generated by the coils allows more localized stimulations, deeper penetration, depth control, and complex nerve stimulation patterns. 
     The controller  120   b  provides signals to the multiplexers  120   a  for selection of the column and row of the coil array  110 . The controller  120   b  controls the power management circuitry  120   c  for providing a value for a particular coil driver circuitry  120   f  shown in  FIG. 1C . The value of Dn determines the amplitude of the current pulse. The timing of the application of Vss and Vcc to a coil driver (see  FIG. 1C ) by the power management circuitry  120   c  is also controlled by the controller  120   b  and determines the phase, duty cycle, and/or frequency of the current pulses provided by the driver circuitry  120   f  to the coil  110   a.    
     Referring now to  FIG. 1C , to program a particular coil, Vcc and Vss are set to zero. A particular column of the array  110  is activated by applying a voltage to Gn and a value Dn is applied to the transistor  127  to set the pulse amplitude value in the gate capacitors  126  for the coil  110   a . The proper bias voltage is set on Dn for the appropriate amount of current that the coil requires, which may be positive or negative. 
     To activate the coil  110   a , Gn and Dn are disabled and Vss and Vcc, e.g. Vss=−5V and Vcc=+5V, are applied to the driver circuit  120   f  for a duration commensurate with the stimulation parameters. Bipolar operation is enabled by connecting the pair of capacitors  126  to the complementary pair of transistors  128 . The coils  110   a  in the array are addressed by a TFT backplane similarly to the way that liquid crystals are addressed by a TFT backplane in a display, utilizing gate and data line multiplexers. The onboard power supply is capable of providing the full range of positive and negative bias voltages for the array, and the controller provides the signals required for activating individual “pixels” in the array and for activating the supply rails. 
     The current pulse through the coils  110   a  generates a magnetic field. Referring now to  FIG. 1D , which shows a single loop coil, and Equation (1), the magnetic field created by each loop, H z   LOOP , increases with the radius of the loop, R, and the intensity of the current, i(t), and decreases with the distance, z, along the axis. 
     
       
         
           
             
               
                 
                   
                     
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     The magnetic field of a coil depends on the number of turns, N, the length, l, the pitch, α, and the amplitude of the current as indicated by  FIG. 1E , which shows a small radius coil, and Equation (2). By increasing N (or the inductance of the coils), and l, we can considerably increase the resulting magnetic field even for a fixed low current and small radius coil, as deduced from Equation (1). Thus, arranging a large number of small coils in an array configuration, as shown in  FIG. 1E  will yield an increase in the magnetic field intensity and penetration to specific regions. 
     
       
         
           
             
               
                 
                   
                     
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       FIG. 1F  is a schematic representation of beam focusing and beam steering within a region of interest  105 . As illustrated in  FIG. 1F , since each coil  110   a  is driven by a coil current that is independent of other coil currents, it is possible to manipulate the field strength of the electric field at a selected location, such as the focal point  105   a  within region of interest  105 , by approximately adding the linear vectors  165  of the individual fields. Thus, tailored stimulations can be obtained with appropriate coil array designs, by selecting the optimal number of elements, array configuration, driving circuits, and current distribution in the coils. 
       FIG. 2  illustrates an embodiment of the neuromodulation device  200 , wherein the device  200  is disposed on a flexible substrate  230  configured as an implantable nerve cuff. The flexible substrate  230  is configured to at least partially surround a nerve bundle  231  and is sutured in position around the nerve bundle  231  via sutures  233  attached to suture anchors  232  located on the flexible substrate  230 . The device  200  includes electrically conductive coils  210   a  arranged in an array  210  and circuitry, e.g., as described in  FIGS. 1A-1F , coupled to energize the coils  210   a  in the array  210  using current pulses that generate an electromagnetic field. The current pulses provided to the coils  110   a  by the circuitry activates the coils  110   a  and creates a tailored magnetic field  220  within the nerve bundle  230 . The circuitry is configured to control one or more parameters of the current pulses, including at least amplitude and phase of the current pulses, such that the electromagnetic field undergoes constructive and destructive interference that focuses and/or steers the electric field within a region of interest of a patient. 
     In the embodiment shown in  FIG. 2 , the region of interest includes the nerve fascicle  205  and does not include nerve fascicles  206 . The magnetic flux density, electric field strength and/or electric field gradient produced by the array  210  is sufficient to activate one or more neurons within the region of interest to provide neuromodulation therapy. For example, the magnetic flux density, electric field strength, and/or electric field gradient produced by the array  210  may be sufficient to activate the nerve fascicle  205  at a specified depth and/or location within the nerve bundle  231  while not activating other nerve fascicles  206  of the nerve bundle  230 . In some scenarios, the neuromodulation therapy may involve using the array  210  to provide a magnetic flux density, electric field strength, and/or electric field gradient to the nerve fascicle  205  at a sub threshold level that is below the activation threshold of the nerve fascicle while not providing the lower magnetic flux density, electric field strength, and/or electric field gradient to the other nerve fascicles  206  of the nerve bundle  230 . 
     The neuromodulation device  200  includes a power supply  240  that optionally includes a battery. In some embodiments, the power supply  240  comprises at least one energy harvesting component, such as antenna  241 , configured to harvest power from a radio frequency (RF) signal generated by an external device (not shown in  FIG. 2 ). The power supply  240  provides power to the circuitry which uses the harvested power to provide the current pulses to the coils  210   a.    
     The coils  210   a  shown in  FIG. 2  are 3D coils. In some embodiments, the device may use coils that are flat planar coils  310   a  disposed on the substrate as shown in  FIG. 3 . In some embodiments, the coils  210   a  are three dimensional coils comprising an out-of-plane micro-structure. The three dimensional coils shown in  FIG. 4 , for example, can be used for the array and enable a magnetic field in a direction parallel to the substrate plane without requiring high aspect ratio processing. 
     The scanning electron micrograph in  FIG. 4  shows an out-of-plane micro-coil in accordance with some embodiments. The coil windings may be made using stress engineered thin films that are deposited by sputtering. The film is photolithographically patterned into strips of micro-springs or elastic members that are subsequently released from their underlying substrate. Upon release, a built-in stress gradient causes the elastic members to curl and form three-dimensional out-of-plane loops that make up the inductor coil. In the coil shown in  FIG. 4 , the free end each of each member contacts an adjacent member. This allows for the formation of a continuous inductor consisting of multiple turns without interruption of the spring metal. To protect the inductor in actual use on a chip or circuit board, the loops can be enclosed in a molding compound. These coils can be mass-produced on prefabricated circuit wafers prior to dicing, bonding, and packaging. 
     In some embodiments, the coils are formed from stress-engineered molybdenum-chromium (MoCr) thin films. The MoCr films are sputter-deposited with a built-in stress gradient so that, when patterned and released from their substrate, they curl into a designed radius of curvature. These micro-machined springs self-assemble 3D scaffolds that are then electroplated with copper to form highly conductive coil windings. The coil arrays can be integrated onto silicon die that also include other circuit elements. Many refractory metals have a common property of acquiring tensile stress when sputtered at high pressures and compressive stress when sputtered at low pressures. This results in a stress gradient that can be induced by changing the ambient pressure during film deposition. A film that is compressive at the bottom and tensile on the surface is, for example, realized by increasing the pressure during sputtering. Pressure control may be accomplished by flowing argon while widening or narrowing an orifice opening to the pump.  FIG. 5A  shows a measured stress versus sputter pressure plot for MoCr. 
       FIGS. 5B and 5C  show a process of forming an out-of-plane coil structure in which two half loops are closed in mid-air forming a loop winding in accordance with some embodiments. A layer of release material  530  is deposited on substrate  500  (for sequential release, two different release layers formed of different release materials may be deposited). Then a layer of an elastic material is deposited on top of the release layer  530 . The elastic layer is photolithographically patterned into a series of individual elastic members  510 - 522 . Each individual elastic member includes a first elastic member (e.g.,  520   a - 520   b ), a contact portion or bridge for connecting between adjacent loop windings (e.g.,  520   b - 520   c ) and a second elastic member (e.g.,  520   c - 520   d ). A layer of solder (e.g.,  520   e ) is optionally formed on the tip of the second elastic member. 
     The loop winding is formed by removing the release window under each first elastic member and each second elastic member. This can be done at the same time, or sequentially, by using a different release material under all the first elastic members than under all the second elastic members. Referring to  FIG. 5C , release of the release layer under the first elastic member  520   a - 520   b  causes a first free portion  520   a  of the first elastic member to be released from the substrate  500 . A first anchor portion  520   b  of the first elastic member remains fixed to the substrate. An intrinsic stress profile in the first elastic member biases the free portion  520   a  away from the substrate  500 . Similarly, release of the release layer under the second elastic member  518   c - 518   d  causes a free portion  518   d  to be released from the substrate  500 . An intrinsic stress profile in the second elastic member biases the free portion  518   d  away from the substrate  500 . A second anchor portion  518   c  remains fixed to the substrate  500 . Pressing and heating causes the solder  518   e  to reflow and join free end  520   a  to free end  518   d.    
     Alternatively, and preferably, the free portions (without solder) can be connected together by electroless plating. Immersion in a plating bath and depositing metal on accessible metal surfaces both thickens all metal lines and creates bridges between proximal surfaces (such as contact portion  520   b - 520   c ). 
     The individual loop halves are shown in  FIG. 5B  as being of approximately the same length. However, the lengths can be varied to aid in the coil formation process. For example, the first elastic members can be made shorter than the second elastic members to ensure that the second elastic members overlap the first elastic members. 
       FIGS. 6A through 6D  illustrate fabrication of the coil array in accordance with some embodiments. As seen from  FIG. 5A , sputter pressures below 2.35 mTorr produce compressive MoCr films, while higher deposition pressures produce tensile films. When patterned and released, such a stress-graded film curls up in a circular trajectory with a radius of curvature given by: 
     
       
         
           
             
               
                 
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     where h is the thickness, Y the biaxial Young&#39;s modulus, and Δσ is the stress difference between the surface and bottom of the film. 
       FIG. 6A  is a cross sectional diagram of layers used to form the coils in the array and  FIG. 6B  is a more detailed view of some of the layers. In this embodiment, fabrication of the coil array starts with the deposition of a metal layer  601 , e.g., a Cu-plated or gold metal layer on top of the substrate  600 . The metal layer  601  serves as a current return path for the coil after it is completed. A 12 to 15 μm-thick low-loss dielectric layer  602  is spin-coated to separate the coil from the substrate and to lower its parasitic capacitance. Vias  610  to the underlying metal layer  601  are then opened. A conductive release/sacrificial layer  603  (not shown in  FIG. 6A  but shown in  FIG. 6B ) may be sputter deposited next, followed by a thin layer of metal  604 , e.g., Au, (also not shown in  FIG. 6A  but shown in  FIG. 6B ), the stress engineered MoCr film  605  already described, and a passivation layer  606 , e.g., an Au passivation layer. This stack is about 1.5 μm thick and the layers  601 - 606  are deposited sequentially in a single pump down. An electrically conductive release layer  603  may be used which is also used as an electrode for electroplating the windings after coil assembly, as will be discussed. The release material has excellent adhesion properties and also functions as an adhesion layer for the stressed films deposited above it. The stress-engineered film  605  is a bi-layer with the first MoCr layer deposited at low pressure on the compressive side of the graph in  FIG. 5A , and the second MoCr layer at a higher pressure on the tensile side of the graph. The resulting built-in stress gradient produces a well-defined mechanical moment on the film. The Au/MoCr/Au  604 / 605 / 606  metal stack is patterned into individual springs that ultimately make up the coil windings. The intrinsic stress of many sputtered thin films depends on the ambient pressure at which the material is deposited. By varying the pressure during sputtering, films can be obtained that are compressively stressed near the substrate-film interface and tensile stressed at the film surface.  FIG. 6B  shows such a stress-graded film  605  sandwiched between two gold layers  604 ,  606 . The stress graded film can be NiZr, MoCr, solder-wettable Ni, or other suitable material. The bottom gold layer  604  forms the outer skin of the coil when released and provides a high conductivity path for electrons at high frequencies. The top gold layer  606  passivates the surface. The metal stack is deposited above a suitable release layer  603  such as Ti, Si, or SiN. The release layer should be a material that can be quickly removed by selective dry or wet undercut etching. Possible etchants for a Si release layer include KOH (wet processing) and XeF2 (dry processing). 
     After defining the release masking windows, the springs are released from the substrate by undercut etching the sacrificial layer. Perforations in the spring metal facilitate the undercut release process. The release mask is designed so that a piece of the photoresist  607  is retained on top of each spring after release. This resist material  607  acts as a relaxable load that retains the springs and prevents the coils from lifting fully during the release process. As shown in  FIG. 6C , when heated, the load layer  607  softens and gradually yields to the built-in stress moment in the spring metal, allowing spring pairs to move in a designed trajectory while self assembling in the air. 
     In some embodiments, the coil structure features an interlocking spring tip that provides a mechanical block that prevents paired springs from curling further after they come together. The interlocking connection of the out-of-plane coil magnified in  FIG. 6D  ensures that the coil diameter is determined by mask design, rather than by the built-in stress. This is desirable because it relaxes the tolerance requirements on the stress profile in the stress-engineered film. Once assembled, the devices are sufficiently robust for handling in and out of plating solutions without coming apart. The assembled structure serves as a three-dimensional scaffold for copper plating. MoCr is a poor electrical conductor, a 5 to 8 μm thick copper skin is electroplated on the scaffold to form low resistance coil windings. The gold finish in the spring metal stack functions as a plating seed. The copper plating not only fills the spring perforations but also electroforms the interlocked seam, joining paired springs to a solid and permanent connection. After plating, the release mask is removed and all remaining release material is cleared. The completed devices are rugged and survive die drops on hard surfaces from heights of over 1 m. This four-mask coil process is compatible with wafer-scale processing and uses conventional sputter deposition techniques, standard photolithography, and simple wet etching techniques. The coils can be seamlessly integrated with other COMES circuitry from single or multiple foundry runs. The intrinsic stress profile in the elastic members discussed above can be designed into a thin film by varying the growth conditions appropriately during deposition to produce coil structures. By adding one or more conductive layers, a coil structure suitable for use as an inductor may be manufactured. 
     The 3D coils discussed in connection with  FIGS. 5 and 6  are metamaterial structures having an intrinsic stress profile. Such metamaterial micron-scale coils allow far greater control of electromagnetic fields as compared with conventional transducer technologies. As understood in the art, the term “metamaterial structure” identifies an artificially engineered structure formed by two or more materials and multiple elements that collectively generate desired electromagnetic properties. A metamaterial structure achieves the desired properties not only from its composition, but also from the exactingly-designed configuration (e.g., the precise shape, geometry, size, orientation and arrangement) of the structural elements formed by the materials of the metamaterial structure. 
     Additional information regarding coils that are suitable for use in the devices and systems discussed herein and their methods of manufacture can be found in commonly owed U.S. Pat. No. 6,646,533 which is incorporated herein by reference. 
     The current pulse injection algorithm used in conjunction with the coils optimizes the current distribution in the coils to enable high-precision targeting and/or to provide tailored stimulus patterns. The current pulse injection algorithm is implements a phased array stimulation wherein coils are selectively energized to provide constructive and/or destructive interference between the electric fields generated by at least some of the coils. The constructive and/or destructive interference in the electric fields allows more localized stimulations, deeper penetration, depth control, and complex nerve stimulation patterns. 
       FIGS. 7A through 7C  illustrate focused magnetic stimulation that is enabled by the devices and systems described herein. FMS can be designed to target and activate nerve fascicles  711 ,  712 ,  713  in specific regions of a nerve bundle  710  without targeting other nerve fascicles  715  in the nerve bundle  710  demonstrating a level of control that is inaccessible to direct stimulation electrode technologies. In  FIGS. 7A-7C , nerve fascicles  711 - 715  have an average diameter of about 150 μm. In  FIG. 7A , the neuromodulation device  700  creates an electric field (E-field) such that nerve fascicle  711  experiences a spatial gradient of the E-field in the x-axis of greater than about 500 V/m 2  which activates nerve fascicle  711  while nerve fascicles  712 ,  713 ,  715  experience spatial gradient of the E-field in the x axis of less than 500 V/m 2 , e.g., about 0 V/m 2  and are not activated. In  FIG. 7B , the neuromodulation device  700  creates an electric field (E-field) such that nerve fascicle  712  experiences a spatial gradient of the E-field in the x-axis of greater than about 500 V/m 2  which activates nerve fascicle  712  while nerve fascicles  711 ,  713 ,  715  experience spatial gradient of the E-field in the x axis of less than 500 V/m 2 , e.g., about 0 V/m 2  and are not activated. In  FIG. 7C , the neuromodulation device  700  creates an electric field (E-field) such that nerve fascicle  713  experiences a spatial gradient of the E-field in the x-axis of greater than about 500 V/m 2  which activates nerve fascicle  713  while nerve fascicles  711 ,  712 ,  715  experience spatial gradient of the E-field in the x axis of less than 500 V/m 2 , e.g., about 0 V/m 2  and are not activated. 
     The approaches described herein provide for steering and focusing the E-field (E) produced by coil array, e.g., a 4×1 or 2×2 coil array, in a region of interest using constructive and destructive interference. The E-field generated by each coil is controlled by the amplitude and phase of the current pulses that energize the coil. In some implementations, the E field generated by coils constructively interferes to create an area within the region of interest having an electric field that is greater than 200 V/m and less than about 50 V/m. The electric field distribution may have a spot size of about 400 μm. The resolution of the stimulation head of the neuromodulation device is dependent on the coil dimensions, e.g., the resolution is about equal to the diameter of coil×3. 
     The electric field manipulation is achieved by modifying the intensity and relative phase of the currents in each coil in the array. By altering the current intensity and phase of individual coils, the ability to stimulate various depths (depth control) may be achieved, as shown in  FIG. 8 . In this simulation, the algorithm optimizes the current distribution at each targeted focal depth to minimize the stimulation spot size. As shown in  FIG. 8  a spot size down to about 400 μm at a focal depth of about 1 mm is achievable using a phased array of 3D metamaterial coils as disclosed herein. 
     As previously discussed, the neuromodulation device disclosed herein may be used in a system that includes a patient information device configured to obtain information about patient conditions. The patient condition information may be obtained through sensors and/or may be input into the patient information device by the patient or other operator. 
     In some configurations, both the neuromodulation device and the patient information device may be a patient-external devices. For example, the neuromodulation device may be a dermal patch and the patient information device may be a patent-external device that communicates with the neuromodulation device through a wired or wired connection. In other configurations, both devices may be patient-internal, e.g., the neuromodulation device may be disposed on an implantable nerve cuff as illustrated in  FIG. 2  and the patient information device be an implantable diagnostic or therapeutic device, such as a cardiac pacemaker, that wirelessly communicates with the neuromodulation device. 
     In yet other configurations, as shown in  FIGS. 9A and 9B , the system may include a neuromodulation device  910  which is a patient-internal device and a patient-external patient information device  920  configured to wirelessly communicate with the neuromodulation device. In some embodiments, the patient information device  920  may be attached to the patient&#39;s skin as shown. 
     For example, the neuromodulation device  910  may be installed using an endovascular approach on the patient&#39;s vagus nerve. The patient information device  920  may comprise a sensing/control module that monitors dynamically changing physiological patient conditions, e.g., heart rate, respiration rate, blood pressure, body temperature etc. The patient information device  920  may sense the physiological state of the patient and generate feedback control signals that are communicated wirelessly to the neuromodulation device  910 . In response to the feedback control signals, the neuromodulation device  910  alters one or more parameters of the current pulses that energize the coils. The feedback control of the patient information device  920  may synthesize and analyze both stimulation and sensing data by utilizing self-learning algorithms, and may be configured to adapt in real-time to enhance therapeutic efficacy. 
     In some embodiments, the patient information device measures biological signals such as heart rate (HR), blood pressure (BP), respiratory rate (RR), body temperature, etc., non-invasively. The patient information device develops a dynamic profile of biological signals in response to focused magnetic stimulation of the selected nerve. Spectral analysis of HR, BP, and RR may be performed to evaluate sympathetic and parasympathetic nervous system contributions. Optimal profiles of biological conditions that provide accurate feedback control for the stimulator function may be developed for each type of nerve stimulated. The optimal profiles may be based on data from a patient population or on individual patient responses to stimulation. The patient information device may be configured to adaptively regulate nerve circuits by continuously assessing the response to the stimulus provided by the stimulator module and reacting accordingly. The patient information device may be configured to use stimulation information, e.g., current pulse amplitude levels, duty cycle, frequency and/or phase along with patient information, e.g., sensed biological data, and/or biological data entered by the patient or other operator, e.g., mood or perception of psychological state. The patient information device may analyze the stimulation information and the patient information utilizing self-learning algorithms, e.g., neural algorithms that mimic human brain function, and may modify the stimulation parameters based on the analysis. 
     The foregoing description of various embodiments has been presented for the purposes of illustration and description and not limitation. The embodiments disclosed are not intended to be exhaustive or to limit the possible implementations to the embodiments disclosed. Many modifications and variations are possible in light of the above teaching.