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Insertion of linear 8.4 μm diameter 16 channelcarbon fiber electrode arrays for single unitrecordingsTo cite this article: Paras R Patel et al 2015 J. Neural Eng. 12 046009

 

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Insertion of linear 8.4μm diameter 16 channelcarbon fiber electrode arrays for singleunit recordings

Paras R Patel1, Kyounghwan Na2, Huanan Zhang3, Takashi D Y Kozai4,7,8,Nicholas A Kotov1,3,5,6, Euisik Yoon1,2 and Cynthia A Chestek1,2

1Department of Biomedical Engineering, College of Engineering, University of Michigan, USA2Department of Electrical Engineering and Computer Science, College of Engineering, University ofMichigan, USA3Department of Chemical Engineering, College of Engineering, University of Michigan, USA4Department of Bioengineering, College of Engineering, University of Pittsburgh, USA5Department of Materials Science & Engineering, College of Engineering, University of Michigan, USA6Department of Macromolecular Science & Engineering, College of Engineering, University ofMichigan, USA7Center for the Neural Basis of Cognition, University of Pittsburgh, USA8McGowan Institute for Regenerative Medicine, University of Pittsburgh, USA

E-mail: cchestek@umich.edu

Received 23 January 2015, revised 24 April 2015Accepted for publication 29 April 2015Published 2 June 2015

AbstractObjective. Single carbon fiber electrodes (d= 8.4 μm) insulated with parylene-c andfunctionalized with PEDOT:pTS have been shown to record single unit activity but manualimplantation of these devices with forceps can be difficult. Without an improvement in theinsertion method any increase in the channel count by fabricating carbon fiber arrays would beimpractical. In this study, we utilize a water soluble coating and structural backbones that allowus to create, implant, and record from fully functionalized arrays of carbon fibers with ∼150 μmpitch. Approach. Two approaches were tested for the insertion of carbon fiber arrays. The firstmethod used a poly(ethylene glycol) (PEG) coating that temporarily stiffened the fibers whileleaving a small portion at the tip exposed. The small exposed portion (500 μm–1 mm) readilypenetrated the brain allowing for an insertion that did not require the handling of each fiber byforceps. The second method involved the fabrication of silicon support structures with individualshanks spaced 150 μm apart. Each shank consisted of a small groove that held an individualcarbon fiber. Main results. Our results showed that the PEG coating allowed for the chronicimplantation of carbon fiber arrays in five rats with unit activity detected at 31 days post-implant.The silicon support structures recorded single unit activity in three acute rat surgeries. In one ofthose surgeries a stacked device with three layers of silicon support structures and carbon fiberswas built and shown to readily insert into the brain with unit activity on select sites. Significance.From these studies we have found that carbon fibers spaced at ∼150 μm readily insert into thebrain. This greatly increases the recording density of chronic neural probes and paves the way foreven higher density devices that have a minimal scarring response.

S Online supplementary data available from stacks.iop.org/JNE/12/046009/mmedia

Keywords: carbon electrodes, neural probes, microelectrodes, MEAs

(Some figures may appear in colour only in the online journal)

Journal of Neural Engineering

J. Neural Eng. 12 (2015) 046009 (18pp) doi:10.1088/1741-2560/12/4/046009

1741-2560/15/046009+18$33.00 © 2015 IOP Publishing Ltd Printed in the UK1

1. Introduction

The ability to record simultaneous neural activity from largeneuronal populations has had a far reaching impact in thefields of neuroscience [1, 2] and neural prosthetics, includingbrain machine interfaces (BMIs) [3–5]. These fields oftenemploy the use of microelectrode arrays (MEAs) which canroutinely record from dozens to several hundred electrodessimultaneously. MEAs are typically composed of siliconshanks, ideally at regular and closely spaced intervals, withone or more electrode recording sites on the individualshanks. Some of the earliest, and still widely used, MEAs aremicrowire arrays consisting of individually placed wires,typically arranged in a linear configuration [1, 6–8]. Othercommonly used high channel MEAs include the Michiganprobe [9–11] and Utah array [12–14]. While extremely pop-ular in their use, all of these MEAs, and their respectiveapplications, would benefit from greater spatial sensitivitywith more closely spaced electrode shanks that can recordfrom more neurons, damage less tissue, and last for decades.

A limiting factor in creating extremely high densityarrays, particularly those that can be implanted long-term, isgliosis. The initial trauma endured by the brain tissue andlocal vasculature upon probe insertion [15–17] can recoverrapidly if the foreign object is either removed [18, 19] ordissolves away [20]. However, when a probe is chronicallyimplanted a persistent reactive response will take place. Thisreactive response is best characterized by a glial sheath orscar, surrounding the probe, composed of activated astrocytesand microglia [18, 21–31]. This scar can extend out to 250 μmfor a typical 15 μm thick silicon shank electrode [18]. Inaddition, the chronic glial sheath contributes to a local zone ofneuronal cell death and neurofilament loss (∼100–200 μmradius) making it difficult to acquire signal [18]. Coupled withthe reactive response is the persistence of leaky vasculature.An implanted neural probe continually disrupts local vascu-lature due to micromotion and mechanical strain [31, 32]leading to the constant release of inflammatory agents[18, 20, 26, 28, 30, 33–35] that perpetuates the local reactiveresponse (see [37] for review). Thus, an MEA with a shankpitch of less than 400 μm would begin to suffer from over-lapping regions of glial scarring, neuronal death, and leakyvasculature, ultimately leading to degraded probe perfor-mance in a continuous scarred region.

The overall shortcomings of traditional MEAs have ledto many groups exploring various design modifications. Somehave tried to develop electrodes with a low Young’s modulusby building the electrodes from ultra-soft composite blendpolymers [36], soft-nanocomposites [38, 39], shape memorypolymers [40], parylene-c [41, 42], polyimide [43], or poly-dimethyl siloxane [44, 45]. These materials have a Young’smodulus that is closer to that of the brain; however, these newprobes introduce new challenges due to bending, buckling,and deflecting of the probe during insertion. In addition, thesepolymer probes use thin-film metal traces which are brittleand geometries that are more susceptible to cracking duringfabrication, assembly, and insertion [46, 47]. To compensatefor the softness of the probe and fragility of the traces, these

softer probes often reinforce their structures by using largerdimensions, which can lead to greater tissue damage and alsogreatly diminishes their ability to be used in a high densityelectrode array [36].

Another approach has been the development of ultra-small probes. Skousen and colleagues showed that a latticestyle electrode, with its greatly reduced surface area, hadsignificantly improved normalized neuronal density (∼90%)within a 50 μm radius, compared to a solid device (∼40%)with the same footprint [48]. This type of electrode designdemonstrates the ability to minimize the scarring response ofthe brain which in turn would allow for MEAs that haveshank pitches much smaller than 400 μm. Many currentdesigns require the use of a shuttle [45, 49] or lattice stylebackbone [48, 50] that can make implantation difficult andpossibly lead to a larger damage radius. In addition, many ofthese smaller devices cannot reach deeper structures in thebrain, can be extremely brittle, and are difficult to achieve inhigh channel counts across large cortical areas compared tostate-of-the-art MEAs.

Ideally an MEA would be made from a soft material thatcan be manufactured with ultrasmall dimensions with suffi-cient strength and durability, and have high channel counts.Unfortunately, such a device does not currently exist due tothe physical limitations and tradeoffs between size andstrength. Our group has taken a different approach by utiliz-ing stiff carbon fiber electrodes, which typically have radii onthe order of 6–7 μm and a Young’s modulus of 241 GPa [51],as compared to silicon’s 165 GPa [52]. Some of the earliestcarbon fiber electrodes were insulated in a pulled glasscapillary tube, which was then back filled with an electrolyte[53]. Once insulated, this electrode can be used as is for acuteelectrophysiology recordings [53]. Coupled with voltammetrymethods the same carbon fiber electrodes can be used todetect dopamine, or other oxidizable neurotransmitters [54],in both animals [55–57] and slice work [58, 59]. Theseapplications are feasible due to the low electrical resistivity ofcarbon fibers. This leads to less thermal noise and enables thefibers to detect small signals without noise contamination[54]. These characteristics have led other groups to developand chronically implant bundled carbon fiber electrodes thatform a hexadectrode bundle, with an overall cross-sectionalwidth of 26 μm [60].

Unfortunately, none of the previously described imple-mentations of carbon fiber electrodes take advantage of all ofthe unique properties inherent to an individual fiber. Thenaturally small cross-sectional footprint of individual carbonfibers leads to less scarring, at a two week time point, and thehigh Young’s modulus makes insertion easier [61]. The highYoung’s modulus would seem to contradict previous find-ings; however, the combined metrics of the small footprintand Young’s modulus give our probes one of the smallestoverall stiffness factors among implantable electrodes whichis well documented by Kozai et al [61]. Instead of the tra-ditional means of insulating the carbon fibers with a glasscapillary, which is brittle and can easily break, our group hasinsulated the devices with parylene-c, an FDA compliant

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material [62]. This allows the fibers to retain their flexibilityor low stiffness while also maintaining their small size.

In this paper we evaluate methods to successfully inserthigh density carbon fiber arrays with site distances ofapproximately 150 μm, which to our knowledge is the smallestrecorded electrode shank pitch for an MEA used in a successfulchronic recording. Initially, the carbon fiber spacing negated ourability to insert each fiber individually without causing damageto neighboring sites. This problem was solved with the encap-sulation of fibers in a temporary poly(ethylene glycol) (PEG)coating. While the PEG encapsulation step was sufficient for theinsertion of arrays with only one or two rows of fibers, it wasdifficult to extend this method to arrays with three or more rowsof fibers, which would allow for greater neuronal samplingwithin a given cortical area. Hence, a second approach using astackable silicon support structure was developed, utilizingdesign cues from the PEG coated arrays, to enable the fabri-cation and insertion of arrays with three or more rows. With thisnew method, single and triple layer devices were fabricated,successfully inserted, and able to record single unit activity inacute surgical preparations.

2. Materials and methods

2.1. Carbon fiber array fabrication and characterization

2.1.1. Array assembly. Custom printed circuit boards(PCBs), designed to fit Tucker-Davis Technologies ZIF

headstages, were manufactured by SpeedyCircuits(Huntington Beach, CA) and designed to create a highdensity array of carbon fibers with a specified site pitch of0.003′′ or 152.4 μm. A Hirose connecter (DF30FC-20DS-0.4V, Hirose, Simi Valley, CA) was soldered to each side ofthe PCB and a shroud placed over the top to conform to ZIFheadstages (Tucker-Davis Technologies, Alachua, FL)(figure 1(a)). Exposed gold coated traces at the tip of thePCB (figure 1(b)) were then individually coated with aconductive silver epoxy (H20E, Epoxy Technology, Billerica,MA) (figure 1(c)) using the tip of a pulled glass capillary. Acarbon fiber bundle (T-650/35 3K, Cytec Thornel, WoodlandPark, NJ) was then separated into individual fibers (d= 6.8 μmaccording to manufacturer specifications [51]), which werethen cut down to lengths of approximately 6–7 mm. A singleuninsulated fiber piece was placed on each of the silver epoxycoated traces, by hand, under a microscope (figure 1(d)). Thisprocess of coating with silver epoxy and placement of thefibers, took approximately 10 min for eight traces. The entireassembly was then placed in an oven to heat cure theconductive silver epoxy at 140 °C for 20 min. These settingsare higher than the manufacturers recommended times, as theoriginal specifications were found to be insufficient for fullycuring the small amount of epoxy used. This same processwas then repeated on the other side of the PCB. A smallamount of insulating epoxy (353NDT, Epoxy Technology,Billerica, MA) was used to protect the trace and carbon fibercontacts (figure 1(e)). A different bake setting (∼120 °C for20 min) was used for the insulating epoxy, as the original

Figure 1. Step-by-step assembly of a carbon fiber array. (a) Bare PCB with attached Hirose connector (light blue arrow) and shroud (whitearrow). Exposed gold traces highlighted in the red box undergo different processing steps which are illustrated in the drawings of (b)–(e).(b) Close up drawing of exposed gold traces from bare PCB. (c) Each trace is coated with a thin layer of silver epoxy using a pulled glasscapillary pipette. Silver epoxy that bridged two traces was removed by wiping the whole board with 70% ethanol. (d) Using forceps, a singlecarbon fiber was placed within each trace. (e) Epoxy is placed on top of the exposed contacts to insulate and protect them from damage.(f) An image of the final product with fibers cut to their desired length.

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specifications were found to be insufficient for fully curingthe small amount of epoxy used. Fibers were cut to desiredlengths (figure 1(f)) using a set of surgical scissors (15003-08,Fine Science Tools, Foster City, CA) and a stereoscopeequipped with a reticle.

To ensure that the initially specified site pitch was wellmaintained after the placement of each fiber, we imaged sixarrays with varying numbers of fibers using a digitalmicroscope (Dino-Lite Pro, AnMo, New Taipei City,Taiwan). Using the distance tool in ImageJ 1.48 v (http://imagej.nih.gov/ij/) the distance between neighboring fiberswas measured. The pitch between fibers was found to be153.7 ± 1 μm (mean ± standard error of the mean, n = 50distances measured).

2.1.2. Parylene-c coating and PEDOT:pTS electrodeposition.Once fully assembled all carbon fiber arrays used in insertiontests and for neural recordings were coated with an 800 nm thickinsulating layer of parylene-c [61] using a Parylene DepositionSystem 2010 (SCS Coatings, Indianapolis, IN). The amount ofparylene-c precursor used was determined by the followingconversion, 2 g of precursor =1 μm thick coating. For thosearrays used to record neural activity the tip of each electrode wascut to expose a small bare carbon fiber site. To lower the exposedcarbon fiber’s site impedance a solution of 0.01M 3,4-ethylenedioxythiophene (483028, Sigma-Aldrich, St. Louis,MO) and 0.1M sodium p-toluenesulfonate (152536, Sigma-Aldrich, St. Louis, MO) was electrodeposited by applying100 pA/channel for 600 s to form a layer of poly(3,4-ethylene-dioxythiophene):sodium p-toluenesulfonate (PEDOT:pTS) [63].All active channels were shorted together during theelectrodeposition step and the total current delivered was scaledaccordingly.

2.1.3. Electrochemical impedance spectroscopy (EIS) andcyclic voltammetry (CV). To verify all deposition andcoating steps EIS and CV measurements were taken with aPGSTAT12 Autolab (EcoChemie, Utrecht, Netherlands),controlled by vendor supplied NOVA software. For allmeasurements, probes were first submerged by 1 mm in a 1xphosphate buffered saline (PBS) solution (BP3994, Fisher,Waltham, MA). A stainless steel rod was used as the counterelectrode and an Ag|AgCl electrode (RE-5B, BASi, WestLafayette, IN) served as the reference electrode. The stainlesssteel rod was chosen as the counter electrode to better matchthe in vivo counter electrode, a stainless steel bone screw(see section 2.2.2). EIS measurements were obtained byapplying a 10 mVrms signal from 10 Hz to 31 kHz [61]. CVmeasurements were obtained by sweeping between 0.8 V to−0.6 V and 0.8 V, a total of three times, at a scan rate of1 V s−1 [61] to rapidly assess the changing properties of theelectrodes as they went through each functionalization step.Custom Matlab (Mathworks, Natick, MA) scripts were usedto determine both the charge storage capacity (CSC) byintegrating the area under the CV curves and frequencyspecific impedance values. It is important to note thatalthough the fibers were inserted 1 mm into the 1x PBS, all

CSC calculations were performed using an area of∼36.3 μm2 as this was the final geometric surface area ofthe carbon fiber’s recording tip. Using this valueconsistently for CSC calculations was done to betterillustrate the changes in electrode characteristics aftereach processing step.

2.1.4. Accelerated soak test setup. Boards with parylene-cand PEDOT:pTS coated carbon fibers had the fiber onlyportion submerged in 1x PBS maintained at 60 °C. At eachtime point the fibers were removed from the heated 1x PBSand rinsed with deionized water. Next the impedance of thefibers at 1 kHz was recorded. Once recordings were completethe assembly was returned to the heated 1x PBS.

According to works by Green et al [63] and Hukins et al[64], equation (1) can be used to determine simulated agingtime that the fibers have undergone:

t t Q10 , (1)TT

37( 37) /10= × −

where t37 is the simulated aging time at 37 °C, tT is theamount of real time that the samples have been kept at theelevated temperature, T , and Q10 is an aging factor that isequal to 2, according to ASTM guidelines for polymer aging[65]. Calculating the simulated time for t    1T = andT    60 C,= ° we find that t    4.92.37 = This value of 4.92 isour acceleration factor and all real time measurements arescaled by this amount to obtain the simulated time.

2.2. PEG facilitated array insertion

2.2.1. Application of PEG coating. The first methoddeveloped to insert carbon fibers arrays, with sufficientlength to reach cortical depths, used 2050MW PEG (295906,Sigma Aldrich, St. Louis, MO), a non-toxic material thattemporarily stiffened the fibers but could be removed with theapplication of sterile room temperature saline or LactatedRinger’s (2B2324, Baxter, Deerfield, IL), via a syringe, overthe course of 15–20 min. Carbon fiber arrays were firstmounted to a micromanipulator and positioned such that thetwo rows of fibers were parallel to the benchtop’s surface.Next, with the aid of a stereoscope, a razor blade covered inaluminum foil was inserted in between the two rows of fibers.The aluminum foil was lightly coated with sterile mineral oil(M5310, Sigma Aldrich, St. Louis, MO) and all excess oilwas removed. The array was then moved such that the toprow of fibers (those visible from a top-down view) weregently resting on top of the oiled aluminum surface. Once thefibers were in position, a coil of wire was wrapped around asoldering iron at 350 °F with a portion of the wire left toprotrude from the tip. The tip of the wire was then used topick up a single flake of 2050 PEG. The PEG was thenallowed to fully melt at the tip of the wire. Once fully meltedthe liquid PEG was applied to the fibers resting on the oiledaluminum surface (figure 2(a)). After cooling the PEG easilyreleased from the oiled surface and the array was retracted.Once clear of the razor blade the array was turned over and

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the same operation was performed on the bottom row offibers.

2.2.2. Surgery for chronic implantation of PEG coated carbonfibers. Chronic implantation of carbon fiber arrays usedadult male Long Evans rats (n = 5) weighing 300–350 g. Ratswere first anesthetized using 5% isoflurane (v/v) to knock-outand then 1–3% isoflurane (v/v) to maintain anesthesia. Thehead was then shaved at and around the area of the incisionsite. The shaved area was triple swabbed using alternatingapplications of betadine and 70% ethanol. Ointment wasapplied to the eyes to keep them from drying during surgery.Once mounted in the ear bars of the stereotax, the shaved areawas swabbed one more time with betadine followed by 70%ethanol. A subcutaneous injection of lidocaine (4 mgmL−1)was given at the proposed site of incision at a maximum

dosage of 4 mg lidocaine per 1 kg of body weight. Afterincision, the skin flaps were pulled away using hemostats andthe skull surface cleaned using a combination of cotton swabsand 2% hydrogen peroxide (v/v). A burr bit (19008-07, FineScience Tools, Foster City, CA) was used to drill seven holesaround the periphery of the skull for seven bone screws(19010-00, Fine Science Tools, Foster City, CA). Next, a2 × 3 mm craniotomy was made over the right hemisphere’smotor cortex using coordinates from a reference atlas [66].Before resecting the dura, a layer of Kwik-Sil (WorldPrecision Instruments, Sarasota, FL) was applied to theskull at the medial and anterior sides of the craniotomy.

Following the resection of the dura, the PEG coatedcarbon fiber array was brought to the surface of the brain. Theexposed fibers were driven into the brain manually using thestereotactic manipulator and depth was monitored using adigital readout. The insertion of the exposed fibers tookapproximately 5–10 s. Once the PEG coated portion of thefibers reached the brain’s surface, manual insertion washalted. Using a syringe filled with room temperature sterileLactated Ringer’s, a small portion of the PEG nearest to thebrain was dissolved away while a surgical spear soaked up thePEG/Ringer’s solution. The fibers were then manually drivenin further, again taking 5–10 s for insertion, and more PEGdissolved away until the final target depth was reached. All ofthe remaining PEG was dissolved away using additionalRinger’s. Additional Kwik-Sil was then applied to the skull atthe lateral and posterior sides of the craniotomy forming abarrier around the craniotomy. Once the Kwik-Sil hadcompletely cured, the craniotomy, exposed fibers, andsurrounding Kwik-Sil barrier were flooded with eitherKwik-Cast (World Precision Instruments, Sarasota, FL) oralginate [67]. Reference and ground wires originating fromthe array’s PCB were attached to the posterior most bonescrew. The PCB was then anchored to all of the skull’s bonescrews using dental acrylic. The skin flaps were brought upover the dental acrylic headcap on each side and suturedtogether at the anterior and posterior ends. Triple antibioticointment was liberally applied around the headcap. Animalswere then removed from the stereotax and allowed torecover on a heated pad placed under their cage. Duringsurgery, animal vitals were monitored using a pulse-oximeter and rectal temperature probe. All procedures andpost-operative care complied with the University ofMichigan’s University Committee on Use and Care ofAnimals.

2.2.3. Electrophysiology recordings and spike sorting.Electrophysiology recordings using chronic implants ofcarbon fiber arrays were done while the rats were awake.Recordings done with acute implants of silicon supportstructures, as discussed below, took place while the rats wereunder ketamine/xylazine anesthesia. All acquisition ofelectrophysiology recordings were taken using a ZC16headstage, RA16PA pre-amplifier and RX5 Pentusa basestation (Tucker-Davis Technologies, Alachua, FL). Duringdata acquisition, the pre-amplifier high pass filtered at 2.2 Hz,

Figure 2. (a) Application of melted 2050 PEG (light blue arrow)using a heated wire (red arrow) while the fibers (yellow arrows) reston a lightly oiled aluminum surface. (b) Carbon fiber array coatedwith solidified PEG (light blue arrows) on both sides withsubmillimeter exposure of the fibers (yellow arrows).

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anti-aliased filtered at 7.5 kHz, and sampled at a rate of∼25 kHz. Each recording session lasted 5 or 10 min.

Recording sessions were imported into Offline Sorter(Plexon, Dallas, TX) and first high-pass filtered (250 Hzcorner, 4th order Butterworth). Each channel was manuallythresholded and the resultant waveforms sorted by a trainedoperator.

2.3. Silicon support structure facilitated insertion

2.3.1. Silicon support structure fabrication. The secondmethod to insert carbon fiber arrays involved the use of acleanroom fabricated silicon support structure. The processstarted with the formation of a trench that eventually serves asa guide for the placement and securement of each carbonfiber. A 10 μm wide and 10 μm deep trench was formed bysilicon deep reactive-ion etching (DRIE) with patternedphotoresist as an etch mask (figure 3(a)). After thoroughorganic/ionic cleaning, the silicon wafer was highly doped(>1019 cm−3) up to ∼12 μm in depth by deep boron dopingand diffusion (figure 3(b)), which is similar to the methodsand dimensions seen in other silicon electrode technologies[9, 68, 69]. The highly doped region was utilized as an etchstop layer of the final ethylene diamine and pyrocatechol(EDP) silicon etch. A dome-shaped cross-section of theeventual individual shanks was achieved by the doping of theexposed trenches. The thickness of doped silicon below thetrench was also ∼12 μm (figure 3(b)) which allowed forsufficient mechanical stiffness during insertion. The shape ofthe shank tips and backend of the silicon guide structure weredefined by another lithography and silicon DRIE step, againusing photoresist as an etch mask. The final step was therelease of structures using an EDP etch step that veryselectively etched the undoped silicon over the highly dopedsilicon (figure 3(c)).

2.3.2. Surgery for acute testing of functional silicon supportstructure. Acute implantation of silicon support structureswith functionalized carbon fibers used adult male LongEvans rats (n = 2) weighing 300–350 g. Rats were first

anesthetized using 5% isoflurane (v/v) to knock-out and thenmaintained under anesthesia with an intraperitoneal (IP)injection of ketamine (50 mg mL−1)/xylazine (5 mg mL−1) ata maximum dosage of 0.125 mL per 100 g of body weight.Regular IP injections of ketamine (50 mg mL−1) were given.The same procedures from section 2.2.2 were then used toprepare the surgical site. Reference and ground wiresoriginating from the probe’s PCB were attached to a singlebone screw.

Following the resection of the dura, the probe wasinserted into the brain to the desired depth by monitoring adigital readout. Electrophysiology recordings were taken andthe animal was sacrificed at the end of the experiment. Duringsurgery, animal vitals were monitored using a pulse-oximeterand rectal temperature probe. All procedures and post-operative care complied with the University of Michigan’sUniversity Committee on Use and Care of Animals.

2.4. SEM Imaging

A FEI Nova 200 Nanolab focused ion beam workstation andscanning electron microscope (FEI, Hillsboro, OR) was usedfor SEM imaging. Prior to imaging, samples were gold sputtercoated with a SPI-Module Sputter Coater (SPI Supplies, WestChester, PA).

2.5. Statistical analysis

A one-sided t-test was used to determine if the dimpling of thebrain for the pointed silicon shanks was significantly less thanthat of the blunt silicon shanks.

ANOVA was used to ascertain if there existed any dif-ference in the alignment of fibers using the two fabricationmethods and varying silicon shank lengths. This was followedby grouped pairwise t-test with a Bonferroni correction todetermine specific group differences.

All calculations were carried out using R 3.0.2 (http://r-project.org/).

Figure 3. Microfabrication steps of silicon support structures. (a) DRIE etch of silicon to form trenches. (b) Boron doping to define overallstructure of the device. (c) EDP etch step to selectively etch silicon and release the structures. (d) Final silicon support structure withoutcarbon fibers.

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3. Results

3.1. Carbon fiber array characterization

3.1.1. Carbon fiber length versus insertion success. Ourprevious work demonstrated the ability to implant individualcarbon fibers, often with the aid of forceps [61]. The153.7 ± 1 μm (mean ± standard error of the mean, n = 50distances measured) pitch of the new carbon fiber arraysmade this previous method impractical during surgery. Toalleviate this problem, fibers were continually shortened. Itwas noted that more fibers would self-insert at shorter lengths.To determine the optimal probe length for reliable insertion,below which isolated fibers would self-insert, one array with14 parylene-c coated fibers was inserted five times (n = 70fibers), for various lengths, into a perfused (PBS only) ratbrain (figure 4). Lengths of 1 mm or greater were found toinsert with a <80% success rate. Lengths of 500 μm werefound to insert with 100% success. These tests were alsocarried out in a 0.6% (w/v) agarose brain phantom [70](figure 4) using one 16 fiber array once per length and one 15fiber array five times per length (n= 91 fibers). It wasobserved during ex vivo tests that the longer fibers did notachieve the same success rates as equal length fibers used inthe agarose test. This is partially attributable to the naturalcurvature of the brain, which caused the fibers to deflect, andthat the agarose phantom is a proxy for the brain and not anexact substitute.

The results from figure 4 were also verified in surgeryusing non-functional parylene-c coated carbon fiber arrays.As expected, the longest fibers (4–5 mm in length) would notinsert and any attempt at manipulation with forceps led tomany fibers breaking off of the PCB. Attempts to insert theshortest fibers were successful as the very tips were able topenetrate the brain, but they could not go any further as thePCB’s back end started to touch the skull’s surface before thefibers reached a depth of 500 μm. At these short lengths(500 μm) and with the current PCB design, reaching layer V/

VI pyramidal neurons in the rat motor cortex, located750–1750 μm below the cortical surface [71], would beimpossible. Also, this target depth does not account for skullthickness [72] and curvature, which necessitates the fibers tobe even longer, making it more difficult for them to self-insert.

To reach relevant cortical depths, while taking into accountfactors such as skull curvature, a coating of 2050 PEG was usedto temporarily encapsulate non-functional parylene-c coatedfibers of various lengths, while leaving the lower ∼500 μmexposed (figure 2(b)). PEG is a safe, biocompatible material,and poses no toxicity to neural tissue [73]. Coating with PEGallowed for a 100% insertion success rate, using three separatearrays with 8, 8, and 6 parylene-c coated fibers one time foreach length (n=22 fibers), in a perfused (PBS only) rat brain forfibers of various lengths (figure 4) as long as the exposed lengthwas <1mm. These results were replicated exactly when using a0.6% (w/v) agarose phantom (data not shown).

3.1.2. Accelerated soak testing of carbon fibers. As a newformulation of PEDOT was used, when compared to ourprevious work [61], an accelerated soak test was carried out todetermine if the new formula could maintain a low impedancethat would allow for chronic recordings over at least a onemonth time course. To test this, fibers were first mounted toPCBs, coated with parylene-c, and then had the tips cut off toexpose a bare carbon site. The exposed fiber tips were thencoated in PEDOT:pTS. Finally, the fully functionalizedcarbon fibers (n= 23 fibers from three arrays) weresubmerged in 1x PBS that was maintained at 60 °C.

Impedance measurements at multiple time points,including just prior to the PEDOT:pTS deposition, weretaken to assess the viability of the PEDOT:pTS coating(figure 5). While the impedance at 1 kHz rose from118 ± 28 kΩ (mean ± standard error of the mean) at simulatedday 0, to 321 ± 56 kΩ (mean ± standard error of the mean) atsimulated day 29.5, these values were found to be acceptable

Figure 4. Insertion success rate (mean ± standard error of the mean)of carbon fiber arrays in agarose without PEG (black, n= 91 fibers),an explanted rat brain without PEG (blue, n = 70 fibers), and anexplanted rat brain with PEG coating (red, n= 22 fibers). PEGcoating enabled a 100% insertion success rate regardless of length orinsertion medium, whereas non-PEG coated arrays suffered from alow success rate at lengths longer than 2 mm.

Figure 5. 1 kHz impedance values (mean ± standard error of themean) of functionalized carbon fibers undergoing an acceleratedsoak test. The pre-PEDOT:pTS impedance value of the fibers withexposed carbon tips (n= 23 fibers) was 6782 ± 355 kΩ. Thisimpedance decreased to 118 ± 28 kΩ, post-PEDOT:pTS depositionat day 0. From day 0 to simulated day 14.8, impedance increased byroughly 100 kΩ, with a similar increase from simulated day 14.8to 29.5.

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for one month recordings. In addition, subsequent chronicimplants also verified the ability of implanted functionalizedfibers to record single unit action potentials, one month postimplant.

3.2. Surgical implantation of PEG coated carbon fibers

After verifying the ability of PEG coated carbon fibers to self-insert and of PEDOT:pTS coated fibers to maintain their low

impedance at 1 kHz, chronic in vivo surgeries were carried outusing carbon fiber arrays with fibers that were approximately4–5 mm in length. During surgery the short exposed ends(∼500 μm) of the PEG coated carbon fibers had no difficultypenetrating the brain. Step-by-step drawings and imagesillustrating the insertion of PEG coated arrays during surgery,along with removal of the PEG can be seen in figure 6. Thissurgical method was used to chronically implant parylene-cand PEDOT:pTS coated carbon fiber arrays (n= 60 fibers

Figure 6. Drawings and surgical images illustrating the insertion process of a PEG coated carbon fiber array. (a) The PEG coated array islowered until the fibers tips are just above the brain. (b) The array is driven down manually over the course of 5–10 s and stopped just beforethe PEG touches the brain’s surface. (c) The craniotomy is filled with room temperature sterile Lactated Ringer’s using a syringe. Any PEGthat comes in to contact with the Ringer’s will start to dissolve. As the Ringer’s becomes saturated with PEG, a surgical spear is used toabsorb the solution and the process is repeated until an additional 500 μm–1 mm of fiber has been exposed. (d) The carbon fiber array ispartially inserted into the brain and more of the fibers have been exposed with the selective removal of PEG from (c). The steps illustrated in(b)–(d) are repeated until the final target depth has been reached which takes approximately 15–20 min. (e) Surgical image of fibers that havebeen partially driven in and PEG that has been partially removed, similar to (d). (f) After reaching the target depth, the remaining PEG isexposed to Ringer’s and is completely removed. (g) An image from surgery showing all fibers inserted into the brain at their final target depthand removal of all remaining PEG, similar to (f). (h) Accessible ZIF connection of a chronically implanted carbon fiber array. PCB wassecured to the skull using dental acrylic.

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from five arrays) into five Long Evans rats. Each processingstep was confirmed with both EIS and CV measurements(figures 7(a) and (c)). EIS measurements are reported for1 kHz (figure 7(b)), in addition to the calculated CSC(figure 7(d)) from the CV curves. All rats showed single unitactivity. A diagram of each probe’s implanted and non-implanted fibers are shown in figure 8(a) as blue and redcircles, respectively. Representative sorted units and high-speed recordings, from days 1 and 31 post-implant, are shownin figures 8(b) and (c), respectively. Full sorted unit and highspeed recording panels for every implanted electrode at days1 and 31 can be seen in figures S1–S4.

The number of successfully implanted carbon fibers(n= 60) is less than the theoretical number of 80 (one 16channel array per animal) due to various factors. The firstcause for missing fibers is that the arrays were partiallydamaged at some point during the fabrication process, startingfrom the initial placement of fibers to the final PEG coating.Second, during implantation, fibers would sometimes stick toone another and need to be separated using forceps or asurgical spear. During this delicate process an unintended

movement could lead to a fiber breaking from the PCB. Third,fibers were purposefully removed if their intended path wasobstructed by a major surface blood vessel. Lastly, the fiberswere actually implanted but their sites were not coated withPEDOT:pTS or in the particular case of one electrode(ZIF_CF_R19, channel 6), day 1 post-operative impedancevalues and electrophysiology recordings indicated a brokenchannel. This last group was classified as broken channels asthey exhibited the same impedance and recording qualities ofthose channels that were actually broken.

3.3. Silicon support structure characterization

3.3.1. Array design and assembly. While the application ofthe PEG coating made insertion of carbon fiber arrayssuccessful, it had some limitations. For carbon fiber arrayswith two layers, a PEG coating, which can be difficult toapply at first, is readily achieved in most laboratories due tothe minimal equipment needed; however, coating with PEGdoes not lend itself well to higher density arrays (i.e., morethan two rows of fibers) using the application method

Figure 7. (a) Impedance magnitude values (mean ± standard error of the mean) were calculated across a range of frequencies at eachfunctionalization step starting with bare fibers, parylene-c coating, exposing carbon fiber tips post parylene-c deposition, and electrochemicaldeposition of PEDOT:pTS (n= 60 fibers from five arrays for each step). Though impedance values for the parylene-c coating step show someperturbations, they likely stem from the excellent insulation properties of this material which makes accurate current flow extremely difficultto measure. Measurements during this step were intended to demonstrate that the coating was present. (b) Impedance magnitude(mean ± standard error of the mean) at 1 kHz for each of the functionalization steps reported in (a). (c) CV plots (mean ± standard error of themean) for each of the functionalization steps reported in (a). (d) CSC values (mean ± standard error of the mean) for each of thefunctionalization steps reported in (a).

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Figure 8. (a) Schematic for each animal’s implanted probe showing broken sites (red) and sites with a successfully implanted fiber (blue).(b) Average waveforms from a single site for each implant in (a) at days 1 and 31 post-implant. (c) Representative filtered (250 Hz corner,4th order Butterworth, high-pass) high speed recordings, displaying sorted unit from (b).

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described earlier. To generate higher density arrays a siliconsupport structure was fabricated that could be stacked with aninterposer.

The silicon support structures were fabricated accordingto the steps outlined in figure 3. Once released from the wafer,the individual devices were mounted on PCBs and a fiber waslaid within the groove of each shank. At the end closest to theprinted circuit board the fibers were secured using a heatcured silver epoxy (H20E, Epoxy Technology, Billerica,MA). The portion of the fibers that lay in the grooves weresecured using an insulating epoxy (301, Epoxy Technology,Billerica, MA) that was also heat cured. The individual siliconshanks had a pitch of 150 μm with a length of either 750 or1000 μm. This fabrication method while not significantlyfaster compared to the arrays using bare traces, has theimportant advantages of even carbon fiber spacing, stack-ability for high density arrays, and a design space that allowsfor the possibility of even closer shanks. The initialparameters of shank length and pitch were largely informedby the previous results of our PEG coated arrays. The pitch ofthe shanks was closely matched to that of the carbon arraysfrom our previous experiments as we already knew that at thisclose distance the fibers could self-insert. In addition, the twosilicon shank lengths were chosen under the assumption thatfor our fibers to both self-insert and reach layer V/VI of rat

motor cortex [71] they would need to extend approximately500 μm beyond the shank ends.

Initial designs had the eight shanks either terminating in ablunt (figures 9(a) and (b)) or pointed end (figures 9(d) and(e)), both of which had a maximal width of ∼30 μm. Initialexperimental surgeries and insertion tests showed the bluntended shanks causing a large amount of dimpling(figure 9(c)). When six blunt silicon shanks of 1000 μmlength from one support structure were inserted to a depth of∼650 μm, brain dimpling was measured at 150.5 ± 41.4 μm(mean ± standard deviation, n = 6). This dimpling was noteasily overcome; however, it was partially alleviated byremoving every other shank. The tradeoff with this fix was the50% reduction of carbon fiber density to a 300 μm pitch,which hinders our ability to create a high density device. Thepointed shanks however, readily inserted into the brain(figure 9(f)). Performing a similar test as before, six pointedsilicon shanks of 1000 μm length from one support structurewere inserted to a depth of ∼650 μm, which resulted in braindimpling of 33.0 ± 11.6 μm (mean ± standard deviation, n = 6).This dimpling was significantly less (p< 0.001, one-sidedt-test) than that of the blunt shanks and allowed us to maintainthe 150 μm pitch needed for our high density arrays.

Another important aspect in designing the support shankswas ensuring that the fibers were as straight as possible, suchthat particular depths could be reached with reasonable

Figure 9. (a) and (b) Photo and SEM image of a silicon support structure with 750 μm long blunt shanks and carbon fibers secured within theshanks. (c) and (f) Insertion of blunt shank support structures resulted in a large amount of dimpling at the brain’s surface (c) as compared tothe pointed shank variant (f). (d) Photo of a silicon support structure with 1000 μm long pointed shanks and carbon fibers secured within theshanks. (e) SEM image of a silicon support structure with a pointed shank.

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certainty. Carbon fibers within shanks of length, 750 and1000 μm, showed a misalignment of 0.208 ± 0.019° (mean±standard error of the mean, n = 69 fibers from nine supportstructure arrays) and 0.157 ± 0.016° (mean ± standard error ofthe mean, n= 71 fibers from nine support structure arrays)respectively, using the edge of the silicon support structure’sbackend as a reference. When compared to devices with noshanks (0.311 ± 0.030°, mean ± standard error of the mean,n= 76 fibers from six arrays), using the edge of the PCBbackend as a reference, both shank lengths showedsignificantly improved alignment (figure 10). All subsequentinsertion tests and experimental surgeries were carried outwith the pointed silicon shanks.

3.3.2. Insertion success. To ensure that the silicon supportstructures could serve as a viable replacement to our PEGcoating, insertion tests with non-functional, parylene-c coateddevices were carried out in 0.6% (w/v) agarose and ex vivousing a perfused (PBS only) rat brain. For each test group asingle eight fiber + shank array (750 or 1000 μm length siliconshanks with 150 μm shank spacing) was inserted five timesfor each length of protruding fiber (n= 40 fibers + shanks pertest). As expected, fibers extending 1 mm or less beyond thesilicon support shanks inserted into the agarose with nearly100% success (figure 11) regardless of shank length. Theinsertion tests into the brain sample (figure 11) were not assuccessful as the agarose tests, however these metrics do lineup well with the previous results seen in figure 4, where theoverall insertion success was lower across all lengths exceptfor that of 500 μm. This again demonstrates that 500 μm ofexposed fiber is the critical length needed for fiber self-insertion.

3.3.3. Functionalization and acute in vivo testing. Given thehigh insertion success rate of the 1000 μm long silicon shanks(150 μm shank spacing) plus 500 μm of exposed carbonfibers, functional tests were carried out to validate thisdesign’s ability to record neuronal signals. As this firstiteration of devices provided no electrical isolation the entiresilicon structure was first coated with an insulating layer ofparylene-c once released from the wafer. Silicon structures

were then secured to PCBs, and fibers laid within theindividual shanks, secured by silver and insulating epoxy(301, Epoxy Technology, Billerica, MA). The entireassembled device was then coated with parylene-c(t= 800 nm) to form an insulating layer around the fibersthemselves and finally the tips of the fibers werefunctionalized with PEDOT:pTS. An electrical connectionbetween the silver epoxy contact point of the carbon fibersand the PCBs exposed traces was made using pieces ofsmall copper wire. Due to the inherent difficulties of placinga wire down for each carbon fiber, only every other shankreceived a carbon fiber (figure 12(a)). Each processingstep was confirmed using both EIS and CV measurements(data not shown) that corresponded well to results shown infigure 6(a).

To demonstrate feasibility of this new device, acute surgeryon two Long Evans rats was performed under ketamine/xylazineanesthesia wherein the probes were implanted in the M1 region ata depth of ∼1.5mm. In the first animal (ZIF_CF_Si_R5) a singleprobe, with four functionalized carbon fibers, was implanted inthe left M1 region. In the second animal (ZIF_CF_Si_R11) twodifferent probes, each with three functionalized carbon fibers,were implanted, one in each hemisphere with M1 as the targetregion. A total of 16 distinct units were detected across nine ofthe ten functionalized channels using manual sorting(figures 12(b) and (c)). The average amplitude for all 16 unitswas 245.9±122.2 μV (mean± standard deviation). These resultsdemonstrate the in vivo capability of our new silicon supportstructures to record high quality, single unit action potentials in arepeatable manner.

3.3.4. 3D support structure—functionalization and acute in vivotesting. To fully validate the ability to insert our 3D designof the silicon supports, a stacked non-functional device wasfabricated. Three silicon supports with 750 μm long shanks

Figure 10. Silicon shank devices show significantly less misalign-ment (pairwise t-test, Bonferroni correction) than devices fabricatedwithout silicon support structures. * indicates significance atp< 0.01, ** indicates significance at p < 0.001. Figure 11. Insertion success rate (mean ± standard error of the mean)

of silicon shanks with various length carbon fibers. Both 750 μm(black, n= 40 fibers + shanks) and 1000 μm (green, n= 40 fibers +shanks) shank lengths showed high insertion percentages in agaroseregardless of fiber length. Tests in an explanted brain showed similarsuccess for 750 μm (blue, n= 40 fibers + shanks) and 1000 μm (red,n = 40 fibers + shanks) shank lengths when fibers were ⩽750 μm. Allresults, regardless of shank length or insertion medium, showed100% success when fibers were 500 μm in length.

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(150 μm shank spacing) and 500 μm of carbon fiber extendingpast the shanks were stacked with an interposer layer of glass(t= 100 μm) (figure 13(a)). This device was successfullyimplanted four times during an acute surgery on a LongEvans rat (supplementary video 1 available at stacks.iop.org/JNE/12/046009/mmedia).

Finally, a fully functionalized device (not shown), similarto that in figure 12(a) was fabricated. In total there were three

silicon layers with a glass interposer (t= 100 μm) betweenlayers. For each silicon support, every other 1000 μm shank(150 μm shank spacing) received a carbon fiber. Due to thelimitations of the PCB, only four of the 12 availablerecordings sites were electrically connected. During an acuteprocedure on one Long Evans rat under ketamine/xylazineanesthesia, recordings were taken with the functionalizeddevice at multiple recordings depths in the M1 region. Singleunit and multiunit action potentials were detected on multiplechannels at various depths (figure 13(b)).

4. Discussion

4.1. Ex vivo characterization and accelerated soak tests

In this work we first sought to characterize the potentialability of carbon fiber arrays to self-penetrate the brain andrecord chronic unit activity. Insertion dynamics and recordingsite stability were evaluated using ex vivo and acceleratedsoak tests, respectively. The ex vivo tests were carried out onPBS only perfused rat brains and showed that 500 μm longcarbon fibers self-inserted with a 100% success rate. Accel-erated soak tests demonstrated an increasing site impedancewhen aged to a simulated time of ∼30 days, but remained lowenough for the detection of single units.

While these tests were not designed to be exact sub-stitutes for chronic in vivo implantation of the carbon fibers,they did serve as methods for rapid, high throughput char-acterization. Additionally, our later in vivo studies confirmedthe findings of these benchtop results.

Figure 12. (a) Image of silicon support structure with carbon fibers.Copper wire (yellow arrow) connects each fiber to traces that leads tothe backend headstage. (b) Representative waveforms detected oneach site across the multiple implants. (c) Representative 100 mssnippet of filtered (250 Hz corner, 4th order Butterworth, high-pass)high speed recordings from each site.

Figure 13. (a) Three stacked layers of silicon supports with carbonfibers forming a 3D device. (b) Average waveforms detected onmultiple sites at various depths from a fully functionalized 3Ddevice.

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Table 1. Commercially available MEAs commonly used in research labs.

MEA technology Shank pitch Electrode shank sizeRecording sitediameter Shank length Materials

Carbon Fibers 150–152.4 μm 8.4 μm diameter with insulation coating 6.8 μm 0.5–5 mm Carbon Fiber + PEDOT:pTSBlackrock Microsystems (Utah Array)[12, 14, 74, 75]

400 μm 90× 90 μm at base, 5 μm tip diameter 40 μm taperingto 5 μm

1–1.5 mm Silicon + Platinum

Microprobes [75] 400 μm 75 μm diameter 2–3 μm 1.5 mm Platinum/IridiumTucker-Davis Technologies [77] 250 μm 50 μm diameter 50 μm 5 mm TungstenNeuroNexus, Inc. [78] 400 μm 50 μm thickness × variable width 15 μm 1.4 mm (implant

depth)Silicon + Iridium

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4.2. PEG

Following the initial carbon fiber characterization usingaccelerated soak tests and PBS perfused rat brains, weestablished the ability to chronically implant individual arraysof 8.4 μm carbon fiber electrodes, with an electrode pitch of∼150 μm. Self-insertion was achieved by taking advantage ofthe carbon fibers’ high Young’s modulus, which allowedthem to penetrate the brain at exposed lengths of 500 μm. Thisunique property, along with a temporary PEG coating allowedus to reach cortical depths of ∼1.5 mm without causing sub-stantial insertion injury as evidenced by the detection ofsingle unit action potentials in chronic implants. This pene-tration depth is sufficient for reaching hippocampus CA1stratum pyramidale layers in mice, cortical layer VI in rats,and cortical layer V in primates. With this preparation wehave avoided the use of an insertion shuttle which canincrease the insertion injury to the tissue immediately sur-rounding the electrode. To the best of our knowledge thesecarbon fiber arrays set a new record for electrode shank pitchwhen compared to the most commonly used chronicallyimplanted functional MEAs (table 1). These higher densityarrays can facilitate the study of horizontal cortical commu-nications, such as lateral inhibition and adaptation, in neu-roscience studies and BMI applications.

The findings from the PEG study suggest that ultra-fineelectrodes made from other materials may also have the rightproperties to self-insert and avoid the use of an insertionshuttle. While the PEG coating method has shown the abilityto chronically implant up to 16 fibers, it is still limited to onlytwo rows of fibers. Increasing the number of rows to three ormore would greatly complicate the existing coating method.To address this challenge, we looked to build upon ourknowledge of carbon fiber insertion dynamics and design aplatform that could support multiple rows of fibers, while alsomaintaining a comparable pitch.

4.3. Silicon support structure

Therefore, a stackable silicon support structure was fabricatedwhich could improve the capabilities and channel count offuture carbon fiber arrays. The blunt ended silicon structureswe originally developed were expected to readily penetratethe brain, given their small footprint of ∼30 (w) ×∼24 μm (h).However, upon testing it was determined that a pitch of150 μm led to a large amount of dimpling. The same styledevice with a tapered end, caused significantly less dimplingat the original 150 μm pitch. This highlights the clear tradeoffbetween a device’s shape, no matter how small, and its pitch.This device was then tested with a single layer of carbonfibers and showed similar insertion success rate versus lengthdynamics as the PEG coated arrays. Further tests using fullyfunctionalized devices demonstrated the ability to recordaction potentials in an acute setting. Lastly, a 3D deviceconsisting of multiple stacked silicon supports was con-structed and was also shown to work in vivo. This new siliconplatform will allow for a modular and highly customizableapproach to building 3D arrays. Depending upon the

application, shank pitch, length, and count can be variedwithin and between individual stacked layers.

4.4. Future directions

While more work is needed to fully understand the long-termviability of these new arrays, the first step is to characterizethe insertion mechanics and assembly methods as thesevariables may impact longitudinal performance. An improvedPCB design with a ground plane between both, layers andtraces, is being actively explored to alleviate any potentialcross-talk between channels. Currently our group is workingon quantifying the long term recording stability of chronicallyimplanted carbon fiber arrays and studying the immuneresponse of such small carbon fiber electrodes through his-tological analysis of multi month animal implants. Histolo-gical studies could help to determine the path and anypossible deviation of the fiber arrays, which may be subject tomicromotion over the duration of the implant.

Additionally, while the goal of this study was not to testinsulation or site coatings, but to develop and evaluateinsertion tools for scalable, high-density carbon fiber arrays,more work needs to be done exploring how chronic implan-tation may affect the performance and adherence of thePEDOT:pTS and parylene-c coatings. The longevity of bothhave been explored by a few groups, some using soak testmethods [62, 63, 79] and others using data from chronicimplants [14, 80], but no definitive conclusions have beenreached. This is partially attributable to the different condi-tions under which each material is used. If current methodsare found to be non-ideal there are a variety of other insu-lating [81, 82] and conductive [83–86] coating methodsavailable (see [87] for review). Lastly, integration of nanos-cale covalent surface chemistries, such as PEG based anti-biofouling coatings [61], neural adhesion molecule coatings[88], and drug eluting coatings [89], are further expected toimprove recording longitudinal performance as well as neu-romodulation or wireless neural stimulation applications [90].

4.5. Conclusion

With the ability for each site to record from a ∼75 μm radius[91], this array theoretically has the potential to record fromall active neurons in a given cortical layer. Longer term,methods need to be developed that will allow for the scalingup of manufacturing throughput and even higher densities ofcarbon fibers. Assembly with silver epoxy could be elimi-nated using a self-aligning anisotropic epoxy that self-selec-tively forms electrical connections between fibers and traces[92]. Innovative methods will be needed for the frontendsilicon support structure but also in the backend where dozensif not hundreds of wires [14] will be needed to transmitelectrophysiological data. Packaging arrays with an ASICcapable of multiplexing hundreds of channels [93] wouldalleviate many wiring issues [2, 94] or completely eliminatethem [95–97]. The convergence of these technologies couldlead to an electrode pitch dense enough to record from allneurons in a given space for long periods of time.

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Acknowledgments

The authors would like to thank Dr Joshua Berke and JeffreyPettibone for helpful discussions regarding the targeting ofsubcortical structures. The authors would also like to thankKaren Schroeder for discussions related to the PEG coatingprocess and Matthew Robbins who assisted with some of theanimal recording sessions. Thanks to the University ofMichigan’s Electron Microbeam Analysis Laboratory and theLurie Nanofabrication Facility for use of their equipment.This work was financially supported by the Department ofEnergy (DE-SC0000957), the National Science Foundation(DMR-0320740), and the National Institute of NeurologicalDisorders and Stroke (1RC1NS068396-0110).

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