Abstract:
An implant for deep brain stimulation (DBS) has an array of electromagnetic microcoils dispersed over the length of the implant. The microcoils produce magnetic fields that are directed into, and induce current in, the adjacent brain tissue. The microcoils may be selectively operated to direct and focus electrical stimulation to targeted areas of the brain. The implant is useful in studying or treating neurophysiological conditions associated with the deep regions of the brain such as Parkinson&#39;s disease, drug addiction, and depression.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application is based on, fully incorporates herein by reference, and claims the benefit of U.S. Provisional Application Ser. No. 61/042,070, filed on Apr. 3, 2008, and entitled “DEEP BRAIN STIMULATION IMPLANT WITH MICRO MAGNETIC STIMULATION ARRAY”. 
    
    
     STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH 
     This invention was made with government support under HD044425 awarded by the National Institutes of Health. The government has certain rights in the invention. 
    
    
     BACKGROUND OF THE INVENTION 
     The present invention relates to deep brain stimulation (DBS) systems to treat certain medical conditions and, more particularly, to an implantable device having an array of selectively operable electromagnetic microcoils to induce currents in surrounding brain tissue. 
     Over 50,000 Americans are diagnosed with Parkinson&#39;s disease each year, with more than half a million Americans affected at any given time. Conventional treatments include pharmaceutical agents that produce dopamine, a neurotransmitter, in an attempt to replenish the low levels found in the brains of those suffering from the disease. Approximately ten percent of people with Parkinson&#39;s disease initially treated with pharmaceutical agents have little to no response. Electrical stimulation of the brain presents an alternative treatment option. 
     Specifically, deep brain stimulation has been used to effectively treat the symptoms of Parkinson&#39;s disease including rigidity, slowed movements, tremors, and walking difficulties. DBS treatment involves the surgical implantation of an electrical stimulator, often referred to as an electrode, lead, or implant, in the basal ganglia. Depending on the observed symptoms and treatment plan, DBS implants may be used to provide unilateral or bilateral simulation in the subthalamic nucleus (STN) or in the globus pallidus internus (GPi). 
     Existing DBS systems include one or more implants having a limited number of electrodes, a programmable current or voltage pulse generator, a battery, and electrical leads. Electrical impulses are created by the pulse generator, directed to the implants via the electrical leads, and continuously delivered to the STN or GPi brain sites via the electrodes up to twenty four hours per day. 
     The technology associated with DBS systems has the potential to help people afflicted with other physical ailments shown to respond to electrical stimulation. For example, stimulation of the brain&#39;s motor cortical areas has been used to help ischemic stroke survivors regain partial use of a weakened hand or arm. Further, it has been suggested that cortical brain stimulation can be successfully used to treat epilepsy. Other neurological disorders may also be treated with stimulation outside of the brain. 
     In spite of the clinical and potential successes, existing deep brain stimulation systems have a number of drawbacks. One drawback is the poor spatial resolution of existing DBS implants. Conventional cylindrical DBS implants have a very limited number of electrodes per implant because of spatial requirements between electrodes to prevent electrophoresis. Because of these gaps between electrodes, electrical stimulation of the brain may not be optimized. Further, during placement and setup of a DBS implant, a large variability exists between the location and size of the stimulation area within the brain and the amount of current to be delivered. Although numerical tools have been developed to estimate the volume of tissue activated (VTA) by each electrode, each DBS implant must still be positioned and set up on a case-by-case basis. 
     A second drawback of existing DBS systems is the large size requirement for electrodes in order to limit the effects of high current densities and electrophoresis. One conventional DBS implant includes cylindrical electrodes with a radius of 0.5 millimeters and a length of 2.5 millimeters. In practice, the dimension of each electrode is roughly equal to the portion of the brain intended to be stimulated, thus limiting the flexibility for spatially selective stimulation of the brain. 
     A third drawback of existing DBS systems is the use of copper-containing electrical leads between the pulse generator and the electrodes. These leads are not compatible with magnetic resonance imaging (MRI) procedures and special precautions must be adhered to during MRI procedures. While copper is not a ferromagnetic material and thus, the electrodes do not move or become dislodged when subjected to strong magnetic fields, large electrical currents may nonetheless be induced resulting in thermal damage to the brain tissue. DBS implants with elongated configurations or that are electronically activated are particularly prone to having induced currents. 
     Efforts have been made to overcome these and other drawbacks of existing DBS implants. Micro- and nano-electrodes, for example, may overcome the poor spatial stimulation characteristics of existing DBS implants and deliver currents into targeted brain regions to provide a more accurate physiological localization and stimulation. However, these implants still use capacitive coupling to deliver an electric current and thus, do not overcome the problems associated with direct electrode-to-tissue contact. 
     Transcranial magnetic stimulation (TMS), overcomes the problems of direct electrode-to-tissue contact by utilizing a non-invasive treatment. TMS devices utilize Faraday&#39;s law of electromagnetic induction that a changing magnetic field can induce electric current to flow in any conductive structure, including human tissue. TMS devices operate by passing a brief electrical pulse through one or more electrically conductive coils positioned adjacent to the human skull. The coils produce magnetic fields at right angles from the coils, through the skull, and into the brain. The magnetic fields, in turn, induce electric fields in the brain tissue to stimulate the nerves. 
     Although non-invasive, TMS has its own drawbacks. Because the intensity of magnetic fields produced by TMS devices decreases very rapidly away from the coil, stimulating deep regions of the brain requires very strong magnetic fields. However, high intensity electric fields (induced by the strong magnetic fields) are known to cause epileptic seizures and other neurological problems. Further, the induced electric fields are not sufficiently focused and, as a result, generalized stimulation throughout the brain may occur. Still further, the amount of electric current used to drive such TMS coils is prohibitively large. 
     Therefore, it would be desirable to have an apparatus that provides effective, accurate, and safe deep brain stimulation. 
     SUMMARY 
     In accordance with one aspect of the present invention, a device for stimulating biological tissue, such as when performing deep brain stimulation, includes an insertable elongated implant with a plurality of microcoils. The implant extends along a longitudinal axis and has a proximal portion and a distal portion. A plurality of electrical conductors extend along the longitudinal axis of the implant and are coupled to the plurality of microcoils. An electrically isolating barrier covers the plurality of microcoils. A coupling connects the microcoils to a power source through the plurality of electrical conductors such that the plurality of microcoils can be driven to produce magnetic fields suitable for deep brain stimulation. 
     In accordance with another aspect of the present invention, a brain stimulation device includes a power source configured to produce a plurality of electric pulses. The brain stimulation device further includes an implant configured to receive the plurality of electrical pulses and generate a magnetic field configured to induce an electrical field adjacent to the implant. 
     In accordance with yet another aspect of the present invention, a deep brain stimulation system includes an implant with a base, an electrical ground layer covering the base, an electrically insulating layer covering the ground layer, a plurality of planar microcoils operable to produce a magnetic field proximate thereto, and a biocompatible dielectric coating covering the plurality of microcoils and the electrically insulating layer. The system also includes a power source configured to power at least one of the plurality of microcoils to produce a magnetic field configured to induce an electrical current in tissue adjacent to the at least one microcoil to perform DBS. 
     The foregoing and other advantages of the invention will appear from the following description. In the description, reference is made to the accompanying drawings which form a part hereof, and in which there is shown by way of illustration a preferred embodiment of the invention. Such embodiment does not necessarily represent the full scope of the invention, however, and reference is made therefore to the claims and herein for interpreting the scope of the invention. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a block diagram illustrating exemplary components of a DBS system including a neurostimulator and an implant with an array of microcoils in accordance with the present invention; 
         FIG. 2  is an enlarged fragmented perspective view illustrating a distal end of the implant of  FIG. 1 ; 
         FIG. 3  is an enlarged plan view corresponding with the encircled region  3  of  FIG. 2 ; and 
         FIG. 4  is an enlarged cross sectional side view taken generally on the line  4 - 4  of  FIG. 3 . 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring initially to  FIGS. 1 and 2 , a deep brain stimulation system  10  includes a stimulator  12  coupled to an implantable device  14  (referred to hereinafter as an ‘implant’) with an array of planar electromagnetic microcoils  16 . The stimulator  12  includes a pulse generator  13  that generates electrical pulses for delivery to a targeted stimulation site in a human brain  18  via the implant  14 . The electrical pulses cause the microcoils  16  to produce magnetic fields that are directed perpendicularly into the brain  18 . The magnetic fields, in turn, induce electrical currents in brain tissue to excite the neurons therein. 
     The implant  14  is configured as an elongated insertion probe with a narrow cylindrical shaft  20  defining a longitudinal axis  21 . The shaft  20  includes a proximal portion  22  with a connector  24  coupled to the stimulator  12  via a pair of leads  26  and a distal portion  28  with a plurality of spaced-apart microcoils  16 . The implant  14  also includes a plurality of electrical conductors  30  ( FIG. 3 ) and a ground layer  32  ( FIG. 4 ) contained within the shaft  20 . Each of the microcoils  16  is coupled to one of the electrical conductors  30  and to the ground layer  32 . 
     Upon implantation of the implant  14 , the microcoils  16  are positioned in close proximity to a target stimulation site for delivery of magnetic field pulses to the brain  18 . Like conventional DBS implants, the implant  14  can either provoke excitation of the brain  18  or momentarily disrupt function of specific cortical regions. Unlike conventional DBS implants that have direct, capacitive contact between metal electrodes and the adjacent brain tissue, the only portion of the implant  14  in direct contact with brain tissue is a biocompatible dielectric sheath, or coating  36  ( FIG. 4 ). Thus, the microcoils  16  are electrically isolated from the surrounding brain tissue and there is no interface therebetween. 
     In one configuration, the implant  14  includes a cylindrical shaft  16  with a diameter of two millimeters (2 mm) and a length of ten centimeters (10 cm). The implant  14  is designed to be surgically placed within a patient&#39;s brain  18  in a manner including, but not limited to, conventional deep brain and cortical electrode implantation techniques as discussed in greater detail below. Each microcoil  16  has an inductance of approximately 13 nH and a stray capacitance of 0.05 pF. 
     The actual number and arrangement of the microcoils  16  about the implant  14  may vary with specific design or application considerations and are considered to be within the scope of the present invention. Other design considerations, such as the geometry (e.g., size, shape, etc.) and placement of the microcoils  16 , may be adjusted depending on the amount or location of neural stimulation for a particular treatment. The induced electric field is a sum of the electric fields induced by each microcoil, and therefore, by changing the driving currents of individual microcoils  16 , the area of neural stimulation can be shaped and targeted. 
     Referring now also to  FIG. 3 , microcoils  16  may be arranged across substantially the entire distal portion  28  of the implant  14 . The microcoils  16  are distributed in an arrayed pattern around the shaft  20 . The microcoils  16  may also be distributed in irregular patterns or have different sizes. Although the illustrated microcoils  16  are spiral-type coils formed with a continuous, multi-turn trace  38  and a have a substantially square footprint, the shape of the microcoils  16  may be circular, oval, rectangular, square, or irregular depending on the particular stimulation requirements. 
     Electrical connections to the microcoils  16  are made at outer bonding pads  40  and inner bonding pads  42  formed at respective ends of the multi-turn trace  38 . The outer bonding pads  40  are connected to a terminal  44  at the proximal end  22  of the shaft  20  via the axially extending electrical conductors  30 . The terminal  44  is coupled to the connector  24  to provide the power to the microcoils  16  via the stimulator  12  and leads  26 . The inner bonding pads  42  are connected to the ground layer  32  by conductive vias  46  ( FIG. 4 ) forming part of the electrical circuit for each microcoil  16 . 
       FIG. 4  is a fragmented cross-sectional view showing the layers and materials of construction of the implant  14 . The implant  14  includes a base  48  and at least four layers applied over the base  48  including, for example, the ground layer  32 , an insulating or dielectric layer  50 , a microcoil layer  52 , and the biocompatible coating  36 . Each of the layers are deposited onto the base  48  which may, for example, be a cylindrical glass fiber, using known deposition methods as described below. 
     The innermost ground layer  32  may, for example, include a three micron (3 μm) thick layer of gold uniformly deposited onto the base  48  using a conventional ion beam deposition method. The ground layer  32  provides a common current return path for each of the microcoils  16 , similar to the ground plane of a typical printed circuit board. 
     The dielectric layer  50  may, for example, include a one hundred micron (100 μm) thick coating of insulating material such as FR-4. The FR-4 material is aerosol deposited over the ground layer  32 . The conductive vias  46  formed within the dielectric layer  50  provide electrically conductive pathways between the microcoils  16  and the ground layer  32 . 
     The microcoil layer  52  includes both the plurality of microcoils  16  and the electrical leads  14  and is situated on top of or at least partially embedded within the dielectric layer  50 . In one configuration, the microcoils  16  are formed from a continuous thin film gold trace  38  and have seven turns. The microcoil  16  may be thirteen microns (13 μm) long by thirteen microns wide (13 μm) by three microns (3 μm) thick in one configuration. In this case, over one hundred microcoils  16  could fit on an implant  14  the same size as a conventional DBS implant. 
     The outermost biocompatible coating  36  may, for example, include a seventy-five micron (75 μm) thick coating of a biocompatible polymeric material, such as parylene. The dielectric, biocompatible material is uniformly applied via chemical vapor deposition at low pressure over the microcoils  16  and dielectric layer  50 . 
     The use of a biocompatible polymeric material, and in particular, parylene, for the coating  36  gives the implant  14  numerous beneficial attributes. Parylene has a low coefficient of friction (e.g., 0.025) such that the implant  14  can be inserted into the brain  18  with minimal damage to adjacent brain tissue. Further, parylene has a low permeability to moisture and gases for example, 0.01% in water), thereby providing stable dielectric properties for the implant  14  over an extended period of time, which is of high importance for brain implants. Still further, parylene exhibits fungus and bacteria resistance, thereby minimizing the likelihood of an immune response. Still further, parylene exhibits high tensile and yield strength (for example, 65,000/6,300 psi), thereby reducing the potential for the coating  36  to be stripped when the implant  14  is inserted into the brain  18 . Further yet, parylene exhibits increased radiation resistance which is beneficial for the sterilization of the implant  14 . Finally, as previously mentioned, parylene has a high dielectric strength (for example, 7,000 V/mil@ 1 mil), thereby providing an effective electrical insulation barrier between the implant  14  and the surrounding brain tissue. 
     The stimulation system  10  may further include a processor  54  to set the amplitude, pulse width, and pulse rate parameters of stimulation pulses based on any of a variety of symptoms or disorders. Although the disclosed stimulator  12  and implant  14  are discussed in the context of a deep brain stimulation system  10  for alleviation of movement disorders such as Parkinson&#39;s disease, other neurological disorders such as epilepsy may beneficially treated with embodiments of the present invention. Likewise, the stimulator  12  may produce stimulation pulses with parameters selected to alleviate chronic pain, gastrointestinal disorders such as gastroparesis or obesity, and pelvic floor disorders such as incontinence, sexual dysfunction, or pain. Accordingly, the implant  14  may be fabricated for stimulation of the spinal cord, gastrointestinal tract, sacral nerves, pudendal nerves, peripheral nerves, and the like. The processor  54  may be realized by one or more microprocessors, digital signal processors (DSPs), Application-Specific Integrated Circuits (ASIC), Field-Programmable Gate Arrays (FPGA), or other equivalent integrated or discrete logic circuitry. 
     The stimulation system  10  may include a switch matrix  56  to apply the stimulation pulses across selected microcoils  16  within a single implant  14  or within two or more implants  14 . The stimulation pulses may be applied in a bipolar or multipolar arrangement, in which multiple microcoils  16  are selected for delivery of stimulation pulses, for example, across or among different microcoil pairs or groups. Alternatively, the stimulator  12  may include multiple pulse generators  13 , each coupled to and controlling a given series of microcoils  16 . 
     A memory  58  may be provided to store instructions for execution by the processor  54  to control the pulse generator  13  and the switch matrix  56 . For example, the memory  58  may be used to store programs defining different sets of stimulation parameters and microcoil combinations. Other information relating to operation of the stimulator  12  may also be stored. The memory  58  may include any form of computer-readable media such as random access memory (RAM), read only memory (ROM), electronically programmable memory (EPROM or EEPROM), flash memory, or any combination thereof. 
     A telemetry unit  60  supporting wireless communication between the stimulator  12  and an external programmer (not shown) may be provided. The processor  54  controls the telemetry unit  60  to receive programming information and send operational information. Programming information may be received from an external clinician programmer or an external patient programmer. The wireless telemetry unit  60  may receive and send information via radio frequency (RF) communication or proximal inductive interaction of a programmer. 
     A power source  62  delivers operating power to the components of the stimulator  12  including the microcoils  16 . The power source  62  may include a rechargeable or nonrechargeable battery or a power generation circuit to produce the operating power. In some embodiments, battery recharging may be accomplished through proximal inductive interaction between an external charger and an inductive charging coil within the stimulator  12 . In other embodiments, operating power may be derived by transcutaneous inductive power generation, e.g., without a battery. 
     Implant Fabrication 
     The fabrication process for the exemplary implant  14  includes a combination of deposition and UV micromolding techniques. In the first step, the glass fiber  48  is placed in an ion beam chamber. Gold is uniformly deposited onto the glass fiber  48  to form the ground layer  32 . Subsequently, a dielectric such as FR-4 is aerosol deposited over the ground layer  32  to create the dielectric layer  50 . Conductive vias  46  are formed in dielectric layer  50 . 
     The microcoils  16  and leads  30  are fabricated in a two step process. First optical lithography is performed by applying, masking, and developing a layer of photoresist on the dielectric layer  50  to form a coil-shaped mold. Second, gold is deposited in the mold. The leads  30  extend over the length of the implant  14  between the microcoils  16  and the terminal  44 . The terminal  44  may be an integral part of a suitable, medical grade connector  24 , such as one produced by the Omnetics Connector Corporation of Minneapolis, Minn. Thereafter, the parylene coating  36  is vacuum deposited over the microcoils  16  and dielectric layer  50 . 
     Computer Simulation of Exemplary Microcoil 
     A theoretical model of a three-turn microcoil  16  was used to calculate the magnetic field generated by such a microcoil  16  and the induced electric field in the surrounding brain tissue using a computer program. The idealized computer model used in these calculations was of a three-turn MEMs inductor coil structure. The theoretical analysis of this model was performed using the computer program Femlab®, a multiphysics modeling software application. Femlab® is a registered trademark of COMSOL AB. 
     The results indicated that microcoils  16  can produce electric fields sufficient to excite brain tissues even when driven by relatively small currents. For example 10 mA. In the simulation, 68,000 elements having 24,776 degrees of freedom, 680 edge elements, and 2,863 boundary elements were used. The microcoil  16  simulation model included a 680×430×600 μm block comprised of three distinct objects including a three turn microcoil made from a series of electrically connected copper traces, a dielectric material surrounding the microcoil, and a tissue substrate. The traces forming the microcoil were modeled with a thickness of 44 μm, a width of 44 μm, and varying lengths. The tissue substrate was modeled as a 680×430×200 μm block located a distance of 100 μm from the copper traces. 
     Femlab® was used to solve the following magnetostatics approximation of the Maxwell equations:
 
−∇·(−σ v ×(∇× A )+∇ V )=0  Eq. (1); and
 
∇×(μ 0   −1 μ r   −1   ∇×A )−σ v ×(∇× A )+σ∇ v =0  Eq. (2);
 
where σ and μ r  are the conductivity and relative permeability {right arrow over (A)} is the magnetic vector potential, and V is the electric potential. The following values were used: (a) copper σ c =1e 6  S/m, μ r =1; (b) dielectric σ c =1e −6  S/m, μ r =1; and (c) tissue σ c =0.3 S/m, μ r =1. The permeability of a vacuum is μ 0 =4*pi*1e−7H/m. All external boundaries were magnetic and electric insulation (i.e., {right arrow over (n)}×{right arrow over (A)}=0 {right arrow over (n)}·{right arrow over (J)}=0), except for the two microcoil boundaries. In these two boundaries, the first that was connected directly to the center of the microcoil was set to magnetic insulation and ground (i.e., {right arrow over (n)}×{right arrow over (A)}=0 V=0) and the other boundary was set to magnetic insulation and 10 mA of inward current flow ({right arrow over (n)}×{right arrow over (A)}=0−{right arrow over (n)}·{right arrow over (J)}=16.10 6  A/m).
 
     The simulations ran for 4,000 seconds on a 3.0 GHz personal computer and the results showed that an electric field having a magnitude of |E|=1.210 5  V/m was induced in the tissue. The simulation further showed current densities with a peak of approximately 50 A/m located directly in the tissue, suggesting that suitable excitation occurs when the microcoils  16  are situated in close proximity to neurons. Current densities in excitable tissue, such as brain tissue, above 10 A/m are known to generate a nerve response or action potential independently from the nerve axon&#39;s size. While the simulations were performed with the three-turn microcoil model, a seven-turn microcoil model may be utilized to have greater induced currents. 
     Computer stimulation may also be used with derived mathematical methods to simulate the effects of size, placement, and number of microcoils  16  in the array on the focality of the stimulation and on the estimated power requirements. 
     Implant Installation and Setup 
     The implant  14  may be positioned and secured into the brain  18  using an MRI system. The first step in the implantation is to non-invasively localize the patient&#39;s STN or GPi regions based on the patient&#39;s anatomical MRI scans. The second step is the functional localization of the STN or GPi sites by recording with microelectrodes at the target nucleus during surgery. The microelectrodes used for recording and stimulation mapping are guided by an MRI-based stereotactic navigation system. The desired location for the target in the STN is in the center of the motor territory. Conversely, the desired target location for the GPi is the anterolateral part of the motor territory 3-4 mm from the internal capsule. The “motor territory” can be localized using electrophysiology. The population of neurons in the STN or GPi with firing rates affected by the patient&#39;s motion (for example, an extremity) is part of the motor territory. 
     Next, the implant  14  is inserted at the site of the microelectrodes. The implant  14  is then set and intraoperative tests are performed to determine appropriate voltage thresholds. After the patient has recovered from surgery, postoperative imaging is used to confirm correct placement of the implant  14 . Finally, permanent programming of the simulator  12  and implant  14  is performed. Importantly, the excitation of neurons in the brain tissue occurs without direct contact, and thus, the heating that may occur in prior art implants at the electrode-to-tissue interface during MR imaging procedures during MR imaging procedures does not occur with the implant  14  with microcoils  16 . 
     The present invention has been described in terms of the various embodiments, and it should be appreciated that many equivalents, alternatives, variations, and modifications, aside from those expressly stated, are possible and within the scope of the invention. Therefore, the invention should not be limited to a particular described embodiment.