referring initially to fig1 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 ( fig3 ) and a ground layer 32 ( fig4 ) 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 ( fig4 ). 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 fig3 , 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 ( fig4 ) forming part of the electrical circuit for each microcoil 16 . fig4 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 . 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 . 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 : ∇×( 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 . 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 .
an implant for deep brain stimulation 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 .