REVIEW OF PERIPHERAL NERVE ELECTRODE TECHNOLOGY Dhonam Pemba Introduction The electrode serves as the critical component of Functional Electrical Stimulation, Brain Machine Interfaces and Neural Prosthesis. The choice of electrode use in any stimulation system would depend on the application. Electrode selectivity varies from measuring surface EMG signals to signal neuron spikes, while invasiveness is inversely proportional to selectivity. Surface electrodes are the least invasive while penetrating intraneural electrodes are the most invasive, accordingly their respective resolution increases with proximity to the neurons. Electrode types can be grouped according to their location; on the surface, epimysial, intramusclar, extraneural, epineural, interfasciular, intraneural. I. Surface Electrodes Surface Electrodes usually consist of stainless steel disks placed on the skin external to the desired muscle to stimulate. The electrode are able to record motor unit action potentials and EMG signals[1–3] and have shown promise in rehabilitation applications[4–7]. Although the electrodes demonstrate capability of detecting muscle activity, their use in for finer motor control would be impractical. The electrodes would require significant current to stimulate muscle for functional motor movement which would cause muscle fatigue and tissue damage. In addition, the large area of the electrodes would produce very low selective stimulation, and accurate muscle recoding for finer control would require significant post signal processing and computation to accurately differentiate motor activity readings[8], [9]. Magnetoneuraphy is an alternative non-invasive method to record magnetic fields in the action in the peripheral nerves[10] but this system would again only be practical for stimulation and depend significantly on post processing[11]. II. Muscle Electrodes To overcome the disabilities of surface electrodes, researchers and clinician have place electrodes directly in contact with the muscle fiber. Methods of muscle stimulation include epimysial electrodes surgically sutured to the muscle[12–14][15], and intramuscular electrodes that are placed inside the muscle[11], [16–19]. The most widespread design of intramsuclar electrodes consist of fine wire inserted into targeted muscle. A novel intramuscular electrode system called the Bionic Neurons(BIONs) was developed to be a self-contained unit. The BIONs are implanted into the muscle with a needled, and are individually addressable, receive power externally and stimulate motor units.[20], [21]. Although epimysial and intramuscular electrodes are able to stimulate and record motor unit action potential, there are several issues preventing success in neuroprosthetic applications. Similar, to surface electrode, the muscle electrodes possess low selectivity due to their distance from the nerve. Even though the electrodes are placed on the muscle, the stimulation actually evokes action potential in the nerve rather thus a higher current would be required stimulate the nerve muscle impedance as opposed to an electrode placed on the nerve directly. In addition, due to the graded recruitment of muscle[22], [23], muscle stimulation would result in muscle fatigue and discomfort[24], [25]. Finally, to achieve complex and fine motor control, electrodes would need to be placed on each individual muscle making the system inconvenient and high maintenance. III. Nerve Electrodes Stimulating the nerve directly has significant advantages of surface, epimysial and intramuscular electrodes. By directly stimulating the nerve, less current can be used while achieve higher selectivity, the electrode can be placed away from the contracting muscle thereby increasing durability, and avert graded recruitment issues to reduce muscle fatigue. Electrodes are either placed outside the epineurium, between epineurium and perineurium, insde the perineurium or through the fascicule. Extraneural Electrodes Extraneural electrodes contact the nerve but are placed outside the epineurium, and including epnineural, helical, book, and cuff electrodes. Epineural electrodes are electrodes that are sutured outside of the nerve. The electrodes are wires, discs or buttons and had success with less complex stimulation needs such as pain therapy or respiration.[26–29] The disadvantages with them are their lower relative selective to other neural electrode and more complicated surgery to ensure stable suture. G.S Brindley pioneered the book electrodes, which have demonstrated success in bladder control by contacting the sacral spinal roots. The device is constructed from implanting electrodes inside channels that are cut out of block biocompatible insulation. The electrode design resembles book, where the gaps between the pages resemble the channels[30–34]. Although the book electrode has had success with spinal root stimulation, the device is too bulky and not designed peripheral nerve stimulation. Recently, a group has attempted to miniaturize the book electrodes for placing individual nerve fascicles between the pages rather than spinal roots[35]. However the miniature book electrode would require the dissection of the nerve into individual fascicles and manual placement in the grooves. Cuff electrodes are arguably the most widely used nerve configuration[36–56], they were designed to improve upon the epineural electrodes by improving the signal to noise and ease of implantation. In general the cuff design composes of an electrode in contact with the epineurium that is encircles the never with biocompatible insulation. Original designs did not utilize micromachining technology and were silicon rubber tubes with a slot[37], blocks of insulating material with a channel[36][57], or flexible insulating material wrapped around a nerve[54], [58].Figure from Loeb et al summarizes in detail the fabrication of several other nerve cuff deigns[59]. With the advancement of thin film micromachining technology, electrode designs have included polyimide as the substrate material, and improved electrode materials and designs[40], [60–65]. Most of the polyimide cuff electrodes are fabricated by depositing and patterning metal between layers of polyimide, and then shaped into a cuff. A critical issue with cuff electrodes is to seal the slot in which the nerve was inserted which was originally done with a silicon elastomer[49], [50], [66],suture[54], [58] or tension[37]. Recent designs have attempted to close the slot by using hinges[45], or advanced polymers with self-coiling abilities[51], [52]. In addition to sheets of biocompatible insulation encircle the nerve, spiral and helical cuff electrodes have been designed[67–73]. Helical electrodes improve fitting while decreasing compression due to self-sizing nature of the spiral. Although cuff electrodes have widely been used, stimulation superficially[74] thus the greatest disadvantages are in the difficulty in selective action of particular fascicles. To improve of selectivity of nerve cuff electrodes, the cuff electrode was redesigned to reshapes the nerve to bring central axon population closer to the electrodes[75], [76]. The flat interface nerve electrode(FINE) were silicon elastomer shaped with a rectangular center opening 1.4 times the cross-sectional area of the nerve(figure [76]). The electrodes have shown effectiveness in controlling gross movements such as knee extension[77], [78] but not complex movements such as ankle control.[79] More recently, Poly (DL-lactic-co-glycolic) acid (PLGA) was added to the a FINE electrode to improve control of closure[80]. The PLGA was attached to a stretched FINE electrode, and upon degradation the FINE electrode deformed to its original equilibrium state. Figure 2.1: FINE electrode (Left), SPINE electrode (Right). Longitudinal electrodes The inventors of the FLAT electrode had previously developed an interfasicular electrode which makes contacts which penetrates the epinerium. The slowly penetrating interfascicular nerve electrode (SPINE) is similar to a traditional cuff electrode but includes blunt electrodes that penetrate epineurm. The SPINE electrode is an interfascicular electrode that is in-between extraneural and intraneural electrodes. Electrodes that penetrate the perineurium are considered intraneural while electrodes outside the epineurium are identified as extraneueral. Intraneural electrodes improve selectivity and signal to noise ratio and reduce cross talk by placing the electrodes inside the nerve and in contact with individual fascicules[81]. Intraneural electrodes possess superior selectivity to extranenural electrodes[74], [82]and have shown the most promise for use in complex motor control neuroprosthesis[83–89]. Restoring hand function is one of the most complex motor movements for a neuroprothesis because high selectivity would be needed to stimulate specific muscles without recruiting others. Success in of stimulating independent specific muscles along with natural hand movements has beech achieved by implanting an intraneural electrode[87]. Intraneural electrodes have shown the ability to detect motor activity in amputee and have helped provide tactile and proprioceptive feedback to discriminate objects without visual cues[90]. By implanting intraneural electrodes, stimulation was able to evoke discrete tactile and proprioceptive sensation in a missing limb.[84], [85] Longitudinally Implanted Intrafascicular electrodes(LIFE) were originally thin insulated wires securely inserted under the perenerium which stimulate and record from individual fascicules. Electrode areas were exposed by heating the insulating layer and the penetrating end was electrochemically sharpened with potassium nitrite[91]. To reduce the mechanical mismatch between tissue and electrode, and decrease cellular encapsulation, flexible substrates were used over the original wire[83], [92]. Polymer fiber was first as a flexible alternative to metal wires, polymer-based longitudinal intrafascicular electrode (polyLIFE) were fabricated by sputter deposition of metals over Kevlar fiber, and dip coating of silicone insulation[83], [93]. Although polyLIFEs improved on flexibility over traditional LIFEs, the active sites precision had not improved. Microfabrication technology was implemented in the newest generation of LIFEs that allowed control over electrode geometry. Fabrication of thin film LIFE (tfLIFE) was similar to polyimide cuff electrodes, where lithography based patterning of metal was sandwiched between two polymer layers[92], [94], [95]. LIFEs offered precise stimulation and record from individual fascicules, but to control complex motor movements various fascicules would need to be selectively stimulated. Thus to selectively stimulate specific fascicules within a nerve fiber bundle would require several implanted LIFEs. Transverse Penetrating Selective stimulation of various fascicules within a nerve fiber bundle would be impractical with LIFEs and unachievable with extraneural electrodes, therefore the best option for electrodes for complex motor function controlled would be transversely penetrating probes. a. Wires Hodgkin and Huxley first demonstrated that action potentials in a squid axon could be detected with a simple wire inserted into a micro glass pipette[96], [97].The basic microwires are easily fabricated and has demonstrated success in neural activity from cortex and peripheral nerves [98–103]. Wire based microelectrodes are constructed from small diameter wires that are then electrochemical sharpened, and finally insulated[99], [104]. Although, these microelectrodes have had success recording neural activity, the fact that selective and independent activation of different fasciculus would require several electrodes limits its application in use with neuroprothesis. b. MEMS Penetrating Electrodes The advancement of MEMS has allowed for microfabrication precision and repeatability to transfer to neural electrodes construction. By incorporating microfabrication technology such as ebeam deposition, sputtering, metal etching and lift off, MEMS based penetrating electrodes now have the capabilities to precisely define electrode geometry as well as a whole larger choice of materials. Bulk and surface micromaching capabilities allow for precision probe shape design and thin film technology allow a larger choice of dielectrics and insulation. MEMS based neural probes were pioneered by University of Utah and University of Michigan. Integrated Circuit technology was first implemented at Stanford by Dr. Wise[105], but Wise moved to University of Michigan and developed with Najafi a planer probe, known famously as the Michigan probe[106], [107]. Along the same time as the Michigan probe was developed, Dr. Norman at the University of Utah produced a three dimensional probe, known as the Utah Electrode Array(UEA)[108], [109]. The Utah Electrode Dr. Richard Norman’s group at the University of Utah fabricated an electrode array referred to as the Utah Electrode Array or UEA[108–116]. The group fabricated a ten by ten electrode array with overall dimensions of 4.2 mm by 4.2 mm Each electrode is needle shaped and can range from 1 to 1.5 mm in length long, with a 80µm diameter at the base. The needles project to a .2 mm thick glass-silicon composite base, which also serves as insulation, see figure 6d[9]. Glass insulation is poured between grooves that are made with a dicing saw in the backside of a silicon square block. The glass/silicon base is grinded down, and patterned with aluminum. The device is flipped over and columns are diced on the top side, and finally the columns are sharpened and insulated. Insulation on the tips is removed by poking the arrays through aluminum foil and etching the exposed tips. This method has been improved recently with parylene C deposition, and a custom holder for photoresist spinning[113]. Other groups have improved charge capabilities of the UEA by coating with new materials[117] or roughening the tips[118]. Figure 2. 2: Utah Electrode Array- Utah Electrode Device including circuit received in May 2008(Top), Varying Length Utah Electrode(Bottom Left) [119], 10 by 10 Utah Electrode(Bottom Right) [111] An improved design with varying length electrodes has also been design to allow stimulation of different regions of a nerve bundle, see Figure 2.2 [110]The authors have also designed a pneumatically actuated insertion method to minimize tissue damage their system was able to an array of 100 needles electrodes into the cerebral cortex with minimal by using a minimum insertion speed for safe entry was 8.3m/s for a depth of 1.5 mm into a cat cortex[109]. There are certain draw backs with the UEA, first of all the fabrication is tedious due to the numerous manual dicing steps required to create one array. Preferably, a more lithography based approach would provide The UEA lacks on chip circuitry thus requiring lead wires that bulk up the size of the device. As can be seen in Figure 3, the electrode arrays are tiny in comparison to the whole device. The circuit board, itself is over 8cm long; the size of this device makes long term implantation impractical. In addition, its structural integrity is in question because it is a composite material and thus introducing an interfacial region. One additional issue is that glass has a lower service temperature than the silicon, which might cause processing obstructions. The Michigan Probe At Michigan Center for Neural Communication Technology planer iridium stimulating arrays were fabricated using surface micromachining and thin-film technology[119–122]. The probes are fabricated by selectively diffusing boron to define the shape of the probes. The thickness of the devices is defined by the thickness of the boron layer. A three layer stack of silicon dioxide, silicon nitride and another layer of silicon dioxide, serve as dielectrics and insulation of the interconnects. Phosophorous doped polysilicon is patterned and serves as interconnects, which is then covered with the three layer stack of dielectrics once more. Next bond pads are made with gold and chromium for adhesion, while electrode sites are made with iridium with titanium for adhesion. The device is finally released by using anisotropic silicon etch composed of ethylenediamine, pyrocatechol, and water.(EDP) Integrated ribbon cables as thin as 2 µm and as thick as 20, as well as a sharper tip angles of less than 10 degrees can also be fabricated using shallow boron diffusion[122]. Later improvement of the original cables was to use polyimide which provides more flexibility and improves durability of the failure prone silicon cables[123]. These probes were able to be integrated with CMOS[119] on chip circuitry. Although planer they can be assembled into three dimensional arrays using orthogonal lead transfers between the planer probes and a platform; Gold electrode plated leads are bent and then the probes are manually inserted into a platform[120]. Despite the great improvements and advancements on the Michigan probe, there are still problems limiting their use in neural implants. Although, the planer geometry is defined by microfabrication capabilities, the depth is dependent on boron diffusion, and therefore limited. In addition, the mechanical mismatch between nerve implantation and these electrodes causes problems of chronic implantation. For long term implantation, a wireless system must be integrated to remove the output leads. There is also the consideration of the tissue-electrode interface. Improvements of the electrodes have been made with bioactive coatings. These techniques have lowered the impedance caused by unfavorable cell activity by coating the electrodes with the conducting polymer polypyrrole (PPy) doped with biomolecules having cell adhesion[124], [125]. Although these systems have improved the interface between electrodes and tissue, there is still the issue of chronic immune response. Preferable drug eluting system must be incorporated to prevent this immune response. Figure 2.3:Michigan probe arrays-Single probes are planer(Top Left) [127], 3 dimensional integration of planer electrodes(Top Right)[128], Cmos Integration(Bottom Left)[120], polyimide cable used instead of silicon ribbon(Bottom Right)[124] Silicon Penetrating Electrodes The UEA and Michigan probe pioneered the field of MEMS based microelectrodes, however the advancements of microfabrication processes and technology brought a whole new era of MEMS based probes. In addition, to the UEA and Michigan probe, researchers have developed various other probes using different materials. In our laboratory, Dr. Jian Wu developed a three dimensional silicon intraneural electrode arrays that utilizes DRIE instead of the dicing saw used in the UEA fabrication[126]. Further improvements by Dr. Wu included the incorporation of a form fitting SU-8 layer to fit the electrode array with peripheral nerves[127]. In addition, Dr. Wu also fabricated a novel curvature-controlled electrode for use in cochlear implants; the probe was made out of polypyrrole which deforms by application of electric field[128]. Researchers in California Institute of Technology utilized emerging parylene technology to develop a more durable alternative to the silicon ribbon and polyimide cables of the Michigan probe[129]. To improve on the uniformity and circuit integration, groups have moved toward dry etching of probe shape. The first probes developed with dry etching were limited to how deep they could etch. Therefore they used a combination of Reactive Ion Etching to define the probe shape, and used backside wet etching to release the probes[130][131]. With the development of deep reactive ion etching (DRIE), thicker probes could be etched. To provide a thicker support structure groups have utilized deep reactive ion etching(DRIE) of SOI wafers[132–134]. DRIE is a dry etching technique and has more control than the wet etching used in the Michigan probe. In addition by using a SOI wafer, the probe thickness is defined by the topside silicon while the support structures depth is limited to the thickness of the SOI wafer. DRIE has also been used by our group to fabricate MEMS based alternative to UEA[126], [135] which requires less manual fabrication with features defined by lithography instead of a dicing saw. Chapter 2 and 3 of this will further optimize our previously fabricated three dimensional electrodes. Figure 2. 4:Silicon probe with parylene wire(Left)[129], Probe made by combination of dry and wet etching (Right) [131]. Polymer Based Probes Silicon based probes provide excellent recording and stimulating capabilities in acute but have not had as parallel success in chronic use. A key issue which degrades electrode charge is the cellular response from the mechanical mismatch between tissue and electrode. Fabrication of flexible neural probes has been a key research issue and has resulted in shift towards polymer based probes. Polymer based probes all have a similar fabrication procedure that involves depositing a sacrificial layer on silicon, then depositing the first polymer, next patterning metal on the first polymer layer, then depositing another polymer layer, exposes the electrodes and finally dissolving the sacrificial layer to release the polymer probe[65], [136–145]. Although, there are a plethora of choices of polymers, polyimide[65], [136], [146], [147], Benzocyclubuten[148–150] and Parylene[143–145], [151], [152] have been used most extensively in neural probes. Along with a more cost effective and easier fabrication procedure, polymer probes can be assembled easier than their silicon counter parts. Takeuchi et al demonstrated three dimensional batch assembly of their polymer probes by using magnetic force[153]. The main disadvantage of flexible probes is that they are not strong enough to penetrate the nerve. But this can be overcome by using a microneedle to first puncture the epineurm and perineum[140] integrate a silicon backbone with the polymer[137] or use a biodegradable polymer to increase mechanical stiff for insertion and later dissolve to increase flexibility[147]. Cellular Response Probes Although flexible probes provide on improvement on long time viability over their silicon counterparts, there still are more promising methods to further improve the lifetime of a chronic implant. The main culprit of electrode failure is due to the host’s systems cellular response. However, the chronic electrode has already by developed over 20 years ago. The neurotrophic electrode consists of gold wires and neurotrophic factors inside of a glass tip[154– 156]. By implanting autologous growth factors into the tip, the electrode integrates with the host tissue by inducing neurite growth, and have achieved viability of 4 years in humans[155]. Although these electrodes have achieved great success in brain computer interface applications[157–160], using them in peripheral nerve neural interfaces will be less successful due to the difficulty in miniaturizing the electrodes and defining active sites. Researchers at the University of Michigan developed a probe to reduce tissue encapsulation around the electrode sites by placing them in a finer lateral area region[161–163]. They were able to increase penetrating strength while placing the electrodes on a smaller area by incorporating an open architecture design. Most recently, the same group utilized parylene-N covered by PEGMA to prevent protein adoption and cellular response[164]. Unlike the neurotrophic electrode, the openarchitecture probes did not contain neurotrophic factors to induce neurite growth, but current research has shown that neurite growth can be inducing by electrical stimulation or neural adhesion molecular coating [165]. Groups have developed electrically stimulated nerve guidance channels to guide nerve growth as an alternative to nerve grafts[166–168]. Although, the guidance channels were not developed for prosthetics, the idea of regenerative electrodes for neuroprostheses were one of the first conceived[169–172]. Figure 2.5: Electrodes design on small surface area to reduce encapsulation[174] Regenerative Electrodes Regenerative Electrodes are another group of penetrating electrodes. They are also known as sieve electrodes and implanted on severed nerves. The electrical activity and/or neurotrophic factors induce never axon growth through the holes in the electrode[40], [171], [173–177]. Although, these electrodes provide the highest selectivity they are limited to use only severed nerves or stumps of amputees. Figure 2.6:Sieve Electrode(Left)[175], Implantation Setup(Right)[177] Fluidic Delivery Probe Figure 2. 7: Various Fluidic Probe Fabrication Techniques-Bonding(Left), Sacrificial layer(Middle), sealing buried channel with dielectrics(Right). The ideal neural probe would have the highest selectivity while evoking the minimal immune response. An electrode with fluidic delivery capabilities could combine the selectivity of penetrating electrodes, the axon regenerative abilities of sieve electrodes and the tissue encapsulation reduction capabilities of neurotopic electrodes. A MEMS based approach would allow for high repeatability, scalability, and precise control over design, while a fluidic channel could delivery either drugs and growth factors to prevent or enhance cellular responses. Figure 2.7 provides an overview of the various approaches to fabricate fluidic channels. University of Michigan researchers have pioneered neural electrode progress, Kensall Wise and David Anderson developed the commercially viable Michigan probe, and their student’s Daryl Kipke and Jamie Hetke continued the legacy and success with NeuroNexus technologies. That legacy still continues today with David Martin leading the field in electrode materials with his startup Biotectix. It would seem fitting that one of the first neural electrodes simultaneous fluidic and electrical capabilities would originate at the University of Michigan[178].Chen et al, fabricated in fluidic channel in the traditional Michigan probes by using shallow boron diffusion to define the channel mask, RIE to etch the channel, EDP to form a continues flow channel by undercutting the mask, and growing and depositing dielectrics to seal the channel[178]. Although, the probes were capable of delivering fluids, they are more practical in research on the cellular response of chemicals due to the extremely shallow depth of the channels. The channels were limited to 15 um due to the channel fabrication and sealing process. In addition, the final release of the probes use EDP as in the Michigan probe fabrication which could potentially destroy CMOS circuitry. To improve CMOS compatibility, UC Berkeley researchers used Deep Reactive Ion Etching of a SOI wafers to both define and release the probes. Their fluidic channel was defined in similar fashion but with a nitride mask, and KOH was to undercut the mask, and low pressure chemical vapor deposition of oxide was used to seal the channels[132]. Figure 2. 8: Silicon Fluidic Probes- Figure Michigan fluidic probe(Left)[180], Berkeley probe (Right)[132]. To improve flexibility, Michigan researchers recently fabricated a parylene C flexible fluidic probe by filling an etched a channel in parylene with sacrificial photoresist and then applying another layer on top to seal the channel[152], [179]. Limitations from this method arise from non-planar deposition of sacrificial photoresist and parylene thickness deposition. The current fluidic probes available at NeuroNexus are made from their Michigan probe mounted to a fluidic tube. Figure 2. 9: Parylene fluidic probe, channel height 5 micrometers[153] Other groups have similarly used sacrificial layers and sealants to construct fluidic neural probes[180–182]. Similarly to the fluidic Michigan probes, sacrificial methods to construct fluidic probes using polyimide or parylene limit the channel depth. Most recently, Kuo et al, inserted a wire into a parylene fluidic probe fabricated from sacrificial method, and thermoformed the electrode into a three dimensional shape with fluidic openings up to 300 um[183]. Boer et al developed the buried channel technology to increase the dimensions of silicon etched microfluid channels. The process involves DRIE of trench, sidewall coverage of trench with either silicon dioxide or LPCVD silicon nitride, removing the coverage on the bottom of trench with RIE, creating buried channel with Isotroptic dry etching, KOH, or HNA, and sealing the channel seal trench with LPCVD of polysilicon dioxide or, silicon nitride[184]. Figure 2. 10:Silicon fluidic probe-DRIE used to make trench (Top), isotropic etching to make channel, and sealed with dielectrics(Bottom).[186] An alternative to sealing the channel with dielectrics or using sacrificial materials is bonding the channel roof to walls of microchannel. Metz et al defined channel walls with polyimide and the channel roof was constructed by bonded and releasing an another polymide layer that was mounted on mylar foil[138].Parker et al used thermocompression bonding Ti foils to fabricate embedded channels[185]. The European Union funded NeuroProbes project utilized wafer bonding to fabricate silicon probes with fluidic channels. Wafer bonding allows for increased channel depth that is now limited to silicon thickness rather than polymer deposition or undercutting mechanics. The NeuroProbes project used DRIE to define channels and the probe shape in one wafer, then mounted wafer to serve as the roof of the channel, the mounted wafer was grinded down to the desired thickness, and DRIE was used to etch out the probe shape on the mounted wafer[186]. The probes then could be assembled into three dimensional configuration with a PDMS based microfluidic connection[187]. Polyimide and parylene probes have the advantage in flexibility, tissue stiffness matching, simplicity of fabrication, while silicon based probes allow for greater control of channel depth and stiff enough for neural penetration. SU8 is a low cost negative resist developed by IBM with single coat spin thickness of 750nm to 450 um[188] and has excellent biocompatibility[189]. Interestingly one of the first uses of SU8 in neural application in 2004[190] and 2005[191] were used as a molds. Park et al fabricated constructed double molds for biodegradable polymer needles by sharpening SU-8 cylindrical posts with RIE[191]. The SU8 molds were used to create PDMS molds which were further used for biodegradable polymers. Kim et al created hollow metallic needles using sacrificial SU8. Tapered SU8 posts were created by backside exposure through glass, which subsequently was covered with metal, then polished to create a hole in posts, and plasma was used to remove the SU8[190]. Lu et al reported one of the first SU8 neural electrodes were in 2006[192] [193], where they fabricated planar SU8 electrodes with grooves for nerves to be inserted. In 2011, Ribeiro et al developed a low cost, simple solution to silicon penetrating electrodes using SU8;They deposited gold in-between two layers of SU8 with omnicoat used to separate the probes from the silicon substrate[194]. In 2012, a Spanish research group under the ERANET-Neuron project fabricated planar SU8 probes under a similar method to Ribeiro et al, with the main difference being the deposition of sacrificial aluminum instead of omnicoat for device separation[195]. In 2009, the Spanish group first demonstrated the viability of SU-8 based neural electrodes with fluidic channels[196], and most recently the same group reported in a different approach to create SU8 embedded microchannels [197]. Their newest involves bonding two patterned SU-8 layers, but crosslinking will increase thermal stability, and glass transition temperature and make reflow and bonding difficult. An alternative to bonding would allow for a simpler process, more accurate alignment, and stronger interface sealing the channels. Summary We presented a review of electrodes for use in neuroprosthetic, brain machine and functional electrical stimulation devices. Intraneural electrodes are necessary to achieve the high selectivity needed to reach central axons of nerve fibers for complex motor control. The tradeoff with increasing selectivity is the increase in invasiveness. Current research understands the need for a chronic electrode to make neuroprosthetic systems a reality. Table 2.1 summarizes the progress made in multi-electrode arrays. Initially, silicon was the gold standard used for neural electrodes, but as researchers attributed the cellular response that degrades the electrode to the mechanical mismatch between silicon and neuronal a trend was pushed to more flexible materials. At the same time, the demand for a probe capable of simultaneous fluidic delivery, and electrical abilities as created because the neurotropic electrode, and others studies demonstrated that neurotrophic factors and drugs can extend the lifespan of probes[154][198], [199][200] while enhance nerve regeneration[201], [202]. Research and progress suggest that the next generation of probes could be made from SU8 as it is biocompatible with customizable flexibility, IC compatibly, and photolithography defined. Table 2.1: Progress in multielectrode array fabrication University/Company Year Material Description Recording and stimulating Recording and stimulating electrode made with EDP and boron diffusion Wire enclosed with neurtrophic factors enclosed in glass cone Stanford University 1970 Si University of Michigan 1985 Si Georgia Institute of Technology 1989 Si/Glass University of Utah 1991 Si University of Michigan 1994 Si University of Michigan 1997 Si/dielectrics University of Southampton 2000 SOI Fraunhofer-Institute 2000 Polyimide University of Michigan 2003 Si/Polyimide University of Tokyo 2003 Arizona State University 2004 Parylene Benzocyclobute ne Arizona State University 2004 Polyimide/SI University of Tokyo Georgia Institute of Technology Emory Universtiy 2004 Parylene 2005 PGA University of Michigan California Institute of Technology 2005 Polyimide 2006 Si/Parylene Universtiy of Texas at Dalas 2006 SU8 Recording and stimulating Silicon Ribbon cables made with shallow diffusion of Boron and EDP etching on non doped Si Fluidic channel made by undercuting mask with wet etchant and seal channel by growing dielectrics Recording and stimulating Recording and stimulating electrodes sandwiched between polyimide Polyimide cables made with same method as Stieglitz et al Recording and stimulating Recording and stimulating Polyimide with silicon backbone tip for stiffness improvement Fluidic probe made with sacrificial layer photoresist SU8 needles used to mold PDMS. PMDS mold used for PGA needle Fluidic probe made by same method as Takeuci et al Parylene Cable made with same methods as Takeuchi et al Reference Wise and Angel[105] Najafi et al[203] Kennedy[154] Campbell et al[108] Hetke et a[122] Chen et al[204] Ensell et al[134] Stieglitz et al[139] Hetke et al[123] Suzuki et al[205] K. Lee et al[150] K. Lee et al[136] Takeuchi et al[144] Park et al[191] Pellinen et al[152] Pang et al[129] Recording and stimulating Lu et al[192] SU8 shank and parylene electrode sites for reduced Seymour et tissue encapsulation al[163] University of Michigan 2006 SU8/ Parylene University of California at Santa Fluidic channel made by thermocompression of two Ti Parker et 2007 Titantium Barbara foil layers al[185] SU8 fluidic probe made by embedded chrome to prevent Fernández et crosslinking al[196] Spain* 2009 SU8 University of Southern Fluidic channel made by thermoforming parylene with California gold wire 2012 Parylene Kuo et al[183] Parylene, SU8 and polyimide coated with PEG for Cheng et stiffness improvement al[147] Polymer coated with PEG 2013 PEG Fluidic channel made by bonding of two patterned SU8 Altuna et Spain** 2013 SU8 layers al[197] *Collaboration between IKERLAN, S. 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