response of brain tissue to chronically implanted neural electrodes

Journal of Neuroscience Methods 148 (2005) 1–18

Invited review

Response of brain tissue to chronically implanted neural electrodes

Vadim S. Polikova, Patrick A. Trescob, William M. Reicherta,∗aDepartment of Biomedical Engineering, Duke University, Durham, NC 27708, USA

bDepartment of Bioengineering, University of Utah, 20 South 2030 East,BPRB 108D, Salt Lake City, UT 84112, USA

Accepted 8 August 2005

Abstract

Chronically implanted recording electrode arrays linked to prosthetics have the potential to make positive impacts on patients suffering from fullor partial paralysis. Such arrays are implanted into the patient’s cortical tissue and record extracellular potentials from nearby neurons, allowingthe information encoded by the neuronal discharges to control external devices. While such systems perform well during acute recordings, theyoften fail to function reliably in clinically relevant chronic settings. Available evidence suggests that a major failure mode of electrode arrays is thebrain tissue reaction against these implants, making the biocompatibility of implanted electrodes a primary concern in device design. This reviewpresents the biological components and time course of the acute and chronic tissue reaction in brain tissue, analyses the brain tissue response of

to enhance

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current electrode systems, and comments on the various material science and bioactive strategies undertaken by electrode designerselectrode performance.© 2005 Elsevier B.V. All rights reserved.

Keywords:Biocompatibility; Chronic; Recording electrode array; Cortical implant; Glial scar; Neuroprosthetic; Wound healing; Foreign body reaction; Neural probe

Contents

1. Introduction. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .2. Immune response to electrode-like implants in the brain. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

2.1. Cells involved in the brain tissue response. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .2.2. Assessing biocompatibility. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .2.3. Mechanical trauma of insertion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .2.4. Long-term inflammation. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .2.5. Glial scar formation. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .2.6. Neuronal response to implant-induced injury. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .2.7. In vivo experimental variability. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

3. Current electrode implant systems. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .3.1. Multiple electrode types. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .3.2. Materials used for the insulating layer. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .3.3. Electrode insertion and implantation procedure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

4. Strategies to minimize the immune response to implanted electrodes. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14.1. Material science strategies. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .4.2. Bioactive molecule strategies. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

5. Perspectives on the current state of the electrode biocompatibility field. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1Acknowledgements. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

∗ Corresponding author. Tel.: +1 919 660 5151; fax: +1 919 660 5362.
E-mail address:[email protected] (W.M. Reichert).
0165-0270/$ – see front matter © 2005 Elsevier B.V. All rights reserved.
doi:10.1016/j.jneumeth.2005.08.015

2 V.S. Polikov et al. / Journal of Neuroscience Methods 148 (2005) 1–18

1. Introduction

Reports that monkeys accurately and reproducibly controlleda robotic arm via chronically implanted cortical electrodes rekin-dled the hope of approximately 200,000 patients suffering fromfull or partial paralysis in the U.S. (Carmena et al., 2003). Theimplants, made out of dozens of wire electrodes, sampled extra-cellular potentials from portions of the monkeys’ cortex (Fig. 1).The potentials recorded from neurons adjacent to the electrodeswere correlated to observed physical motion, eventually allow-ing the researchers to translate the neuronal activity directly intorobotic arm movements. Because of the length of time requiredfor training and the eventual clinical necessity of a long-termbrain–machine interface, the implants were chronic, that is theywere permanently implanted into the monkeys’ brains.

The purpose of these implanted devices is to record signalswith high signal to noise ratios (SNR) from as many individ-ual neurons, termed single units, as possible. Single units mustbe separated from the background electrical noise and from theoverlapping signals, or multi-unit potentials, of other nearbyneurons. Principle component analysis and other signal process-ing techniques can be used to separate single units from noiseand multi-unit potentials (Williams et al., 1999b), with a targetSNR for recordings around 5:1 (Maynard et al., 2000; Rouscheet al., 2001).

A number of engineering groups are developing multi-c ionsT udinm( a-cb euri resed evem s, ap lop-m

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These studies are reminders of the different requirements ofa chronic brain implant suitable for use in basic research andthe reliability threshold needed for clinically relevant neuro-prosthetics. For clinical applications, implanted microelectrodearrays intended as control and communication interfaces needto record unit activity for time periods of the order of decades(Nicolelis and Ribeiro, 2002) and must have a substantially highperformance reliability.

This review highlights what is known of the tissue reactionto implantable electrodes, commonly referred to as the “bio-compatibility” of the implant. In addition, we survey some ofthe potential strategies used to improve the biocompatibility ofchronically implanted cortical electrodes. The purpose of thereview is not to provide an exhaustive examination into any oneresearch thrust, but rather to present a broad view of the issuessurrounding this biocompatibility problem so that computationalneuroscientists, biomedical engineers, materials scientists, andneurobiologists can come together with a common understand-ing of the issues and advance the field.

2. Immune response to electrode-like implants in thebrain

Available studies suggest that the greatest challenge to obtain-ing consistent or stable intracortical recordings is the biologicalresponse that the brain mounts against implanted electrodes.I vadet ), it isn lved.T onset nal-y .(

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hannel recording electrode arrays for chronic applicathese designs span several different technologies, inclicrowires (Williams et al., 1999a; Kralik et al., 2001), polymers

Rousche et al., 2001), and different types of silicon micromhined implants (Drake et al., 1988; Campbell et al., 1991). Theasic sensor and instrumentation requirements of such n

mplant systems are well established, and in general, the pesigns perform as intended in short term studies. Howany designs perform inconsistently in chronic applicationroblem that poses a significant barrier to the clinical deveent of this promising technology.For example,Nicolelis et al. (2003)reported a 40% drop in th

umber of functional electrodes between 1 and 18 months.of 11 electrode shafts in Rousche and Normann’s impla

rray recorded signals at implantation, and the number ofrodes recording signals dropped to 4 of 11 after 5 moRousche and Normann, 1998). In a study byWilliams et al.1999a), eight cats were implanted with microwire electrorrays and electrical activity was recorded over time. Thre

he electrode arrays failed within the first 15 weeks, presbly due to the tissue reaction and loosening of the skulsed to keep the electrode in place. The other electrode aemained active until the cats succumbed to unrelated meomplications between 15 and 25 weeks post-implantationLiut al. (1999)found a large amount of variation in the stabif neural recordings between different electrode shafts oame array and between arrays implanted in different cats.euronal recordings in the study either grew in stability uay 80 post-implantation, after which the recording rematable, or degenerated from a high stability to nearly no siground 60 days post-implantation.

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n order to design recording electrodes that minimize or ehe tissue response of the central nervous system (CNSecessary to understand the biological mechanisms invohe following provides an overview of the CNS tissue resp

o implanted needle-like materials. For a more in-depth asis, the reader is referred toLandis (1994)or Berry et al1999).

.1. Cells involved in the brain tissue response

There are several distinct cell populations involvedhe inflammatory and wound healing response to matemplanted in the CNS, also referred to as the “foreign besponse.” The cell most commonly associated with brain ts the neuron. The location of the neuron soma and its car processes relative to the electrode recording sites deterhe strength and quality of the recorded electrical signal. Tetical models predict that action potentials cannot be obsbove noise farther than approximately 130�m from a record

ng site and that the neuron soma contributes the majority oecorded signal (Eaton and Henriquez, 2005). Direct measureents suggest however, that the maximum distance is muc

omewhere between 50 and 100�m (Mountcastle, 1957; Ra962; Rosenthal, 1972; Henze et al., 2000). Thus, the distancequired to maintain a recording between an electrode sitneuron cell body is of the order of cell dimensions.Although neuronal networks are responsible for informa

rocessing and ultimate control of bodily functions, neuake up less than 25% of the cells in the brain (Purves et al001). The remaining tissue consists of the glial cells (oligodrocytes, astrocytes, and microglia) and vascular-related t

V.S. Polikov et al. / Journal of Neuroscience Methods 148 (2005) 1–18 3

Fig. 1. Schematic description of two potential applications of brain–machine interfaces. (a) A “brain pacemaker” that monitors neural activity to detect seizures.When seizures activity is detected, the implant sends a signal to a nerve cuff electrode or a mini-pump for drug delivery to stop the seizures. (b) Electrode arrayssample the activity of large populations of neurons to control the movements of a prosthetic arm (fromNicolelis, 2001).

Oligodendrocytes are the myelin forming cells of the CNS, whileastrocytes and microglia are the main effectors of the brain’sresponse to injury.

Astrocytes make up 30–65% of the glial cells in the CNS(Nathaniel and Nathaniel, 1981). They contain an elaborate setof cellular extensions, giving them a star-like appearance in his-tological preparations of brain tissue sections and in cell culture.They provide growth cues to neurons during CNS development,

mechanically support the mature neuronal circuits, help controlthe chemical environment of the neurons, buffer the neurotrans-mitters and ions released during neuronal signaling, and evenmodulate the firing activity of neurons (Kimelberg et al., 1993;Kettenmann and Ransom, 1995; Purves et al., 2001). Special-ized astrocytic extensions, termed end feet, abut capillary wallsand aid in the transfer of nutrients across the blood–brain bar-rier. Similar processes weave together to form the 15�m thick


Fig. 2. Schematic representation of astrocyte activation to a reactive phenotype.

glia limitans, the glial boundary between CNS and non-CNSstructures that can be clearly seen even between the CNS andthe peripheral nervous system (PNS) (Nathaniel and Nathaniel,1981). Astrocytes are characterized by 8–10 nm diameter inter-mediate filaments of polymerized glial fibrillary acid protein(GFAP), which is considered an astrocyte specific cell marker

(Eng and DeArmond, 1982). Activation of the astrocytes byinjury transforms the cells into a “reactive” phenotype (Fig. 2)characterized by enhanced migration, proliferation, hypertro-phy, upregulation of GFAP, changes in the number, and distribu-tion of cellular organelles and glycogen deposits, and increasedmatrix production (Landis, 1994). Immunostaining for GFAP

matio
Fig. 3. Schematic representation of microglial transfor n from a resting to an activated, mobile, highly phagocytic cell.


is the most common method for astrocyte identification and fordetermining the extent of “reactive gliosis”, the term used todescribe the activation, hypertrophy, and proliferation of astro-cytes in response to injury (Bignami et al., 1980; Eng andDeArmond, 1982).

Microglia are the other major glial cell type involved in thebrain’s wound healing response, constituting 5–10% of the totalnumber of glial cells in the brain (Ling, 1981). These cellsappear to arrive in the brain via a prenatal infiltration of theCNS by blood-borne hematopoetic cells and remain thereafteras the resident macrophages in neural tissue. Primarily, theyact as cytotoxic cells killing pathogenic organisms or as phago-cytes secreting proteolytic enzymes to degrade cellular debrisand damaged matrix after injury or during regular cell turnover(Streit, 1995; Purves et al., 2001). Microglia reside in an inac-tive or ramified, highly branched state until “activated” via injurymediated mechanisms (Fig. 3). Upon activation, they begin toproliferate, assume a more compact “amoeboid” morphology,phagocytose foreign material, and upregulate the production oflytic enzymes to aid in foreign body degradation (Ling, 1981).When CNS damage severs blood vessels, microglia are indistin-guishable from the blood borne, monocyte-derived macrophagesthat are recruited by the degranulation of platelets and the cel-lular release of cytokines. Macrophage-like cells arriving fromdamaged blood vessels, those present as perivascular residentcells, and microglia seem to perform the same functions inr thisr

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2.2. Assessing biocompatibility

The literature reflects a concerted effort to identify and char-acterize the changing temporal characteristics of the foreignbody response in the CNS. The most common method to assessthe tissue response is to implant model probes into test animalsand sacrifice the animals at various time points to evaluate theprobe and adjacent neural tissue. Direct evidence linking tissueresponse to device performance, however, has not been studiedextensively. The majority of available evidence is derived fromstudies that examine the brain tissue response to passive or non-functional implants. Because biocompatibility is a functionalcharacteristic, which strictly speaking requires an analysis of tis-sue response using an electrically functional implant, the effectof the tissue response on device function is mostly inferred.

In general, electrodes are removed following tissue fixationand the excised tissue surrounding the implant is sectionedand stained to identify cell numbers, locations, types, and by-products. For the most part, the results are subjective, consistingof a description of a particular stain in comparison to adjacent,uninjured tissue from a few selected sections with the number ofanimals per time point and number of sections analysed varyingconsiderably from study to study. A few studies have exam-ined the explanted electrodes with histological methods (Turneret al., 1999; Szarowski et al., 2003; Biran et al., 2005). Most ofthe histological studies have focused on the reactive astrogliosis,a f ther

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esponse to injury and will not be distinguished hereafter ineview.

Microglia are also known to secrete multiple solubleors that affect a variety of processes and signaling pathwhich makes it difficult to understand their precise role inrain tissue response to implanted materials (Banati et al., 1993).hey are a potent source of MCP-1, a chemokine that recacrophages and activated microglia (Babcock et al., 2003),nd of pro-inflammatory cytokines, such as IL-1 (Giulian etl., 1994a), IL-6 (Woodroofe et al., 1991), and TNF-� (Giuliant al., 1994b; Sheng et al., 1995; Chabot et al., 1997). Further-ore, microglia are known to secrete, either constitutivel

n response to pathological stimuli, neurotrophic factorsid in neuronal survival and growth (Nakajima et al., 2001),

ncluding NGF (Elkabes et al., 1996; Nakajima et al., 200),DNF (Nakajima et al., 2001), and NT-3 (Elkabes et al., 1996).inally, microglia produce various cytotoxic and neurotoxic

ors (reviewed byNakajima et al., 2001) which can lead teuronal death in vitro (Giulian et al., 1994a) and in vivo follow-

ng traumatic injury (Giulian et al., 1993a,b). These cytotoxiactors include excitatory amino acids, such as glutamate,ive oxygen intermediates (ROIs) generated as a result oicroglial “respiratory burst” (Auger and Ross, 1992), such asydrogen peroxide (H2O2) or superoxide anion (O2−) (Giuliannd Baker, 1986), and reactive nitrogen intermediates, sucitric oxide (NO) (Zielasek et al., 1996; Minghetti and Le998). It has been suggested that the presence of insolubleials in the brain may lead to a state of “frustrated phagocytor an inability of macrophages to remove the foreign body, wesults in a persistent, constitutive release of neurotoxictances (Weldon et al., 1998).

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nd less so on the microgliotic response or on the effect oesponse on the surrounding neuronal population.

Early studies relied on Hemotoxylin and Eosin (H&E)etermine the location of neurons and other relevant cell telative to the implant (Edell et al., 1992), although H&E is notell-type-specific stain. To gain more specificity, immunostng for cell-type-specific proteins, such as GFAP for astrocnd NeuN or MAP-2 for neurons, is used to localize cells anetermine their state of activation (Menei et al., 1994). Follow-

ng activation, microglial cells exhibit increased expressioertain leukocyte-associated molecules, including CD68gnized by the ED-1 antibody) and Mac-1/CD11b (recogny the OX-42 antibody), or of various lectins (Thomas, 1992). Atain for vimentin, an intermediate filament expressed in rive astrocytes, microglia, and perivascular cells, can alsmployed (Cui et al., 2003) although specificity is lost. To test train’s response to acute injury, as opposed to chronic im

ation, the implant can be removed soon after implantatioerely used to create the initial insertion injury without ac

mplantation (Yuen and Agnew, 1995; Biran et al., 2005).Although staining allows for the identification of react

ells around an implant, other methods have been usxamine the extracellular environment. Microdialysis sampllows the user to assess the chemical environment surrou

he implant without the presence of confounding cells (Khannd Michael, 2003). The dialysis membrane can be adjus

o pass only certain molecular weight compounds, and theected dialysate is then tested (often through an immunoassetermine the presence of different compounds in questionodification of the microdialysis sampling paradigm, hol

ber membranes (HFM) of poly(acrylonitrile-vinyl chlorid


(PAN-PVC) have been implanted to mimic electrodes (Kim etal., 2004). The foreign body response around an HFM implantis then characterized by immunostaining for cells and testingextracellular milieu collected from the hollow inner compart-ment. Finally, others have studied the neural tissue responseto implants using hippocampal slice cultures grown on porousmembranes or on silicon chips with recording capabilities withvariable success (Kunkler and Kraig, 1997; Kristensen et al.,2001).

2.3. Mechanical trauma of insertion

Few studies have examined the initial impact of insertion, andit appears that the first encounter of cortical tissue with a needle-like electrode is a violent one. As the electrode is inserted intothe cortex, its path severs capillaries, extracellular matrix, glialand neuronal cell processes. The electrode may pull and snapextracellular matrix materials as it progresses deeper into tissue,push aside tissue that had once occupied the electrode space,and induce a high-pressure region surrounding the electrode.

This mechanical trauma initiates the CNS wound healingresponse, a response that shares similarities to wound healingresponses of other tissues. Disruption of blood vessels releaseserythrocytes, activates platelets, clotting factors, and the com-plement cascade to aid in macrophage recruitment and initiatetissue rebuilding. Insertion induced accumulation of fluid andn presp acks upoeA –5%o dems essem aveb .

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2.4. Long-term inflammation

The brain tissue response of chronically implanted elec-trodes would be less of an issue if the foreign body responsedisappeared a few weeks after implantation as observed withstab wounds. However, once the acute inflammatory responsedeclines, a chronic foreign body reaction is observed. This reac-tion is characterized by the presence of both reactive astrocytes,which form a glial scar (detailed in the next section), and acti-vated microglia (Stensaas and Stensaas, 1976; Turner et al.,1999; Szarowski et al., 2003; Biran et al., 2005).

In many studies, a significant portion of cells in damagedneural tissue do not stain for GFAP, suggesting the presence oflarge numbers of activated microglia at the surface of implantedbiomaterials long after the initial wound healing response iscomplete, and perhaps as long as the material remains in con-tact with brain tissue (Stensaas and Stensaas, 1976, 1978; Winnet al., 1989; Edell et al., 1992; Menei et al., 1994; Kunklerand Kraig, 1997; Mofid et al., 1997; Emerich et al., 1999;Mokry et al., 2000; Szarowski et al., 2003). For example, nearly25% of electrode tracks in a multi-pronged silicon probe arrayshowed macrophage-like cells present 6 months after implan-tation (Schmidt et al., 1993). Activated microglia will attemptto phagocytose foreign matter for eventual degradation. When25�m polymeric microspheres were implanted into rat cortex,they were all phagocytosed by the activated microglia within 2m er oft da lia int in ar eeks,a th at1 ts ctioni

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ecrotic nervous tissue causes edema, further adding to theure surrounding the implant. When a 10× 10 array of siliconrobes was implanted in feline cortex, 60% of the needle trhowed evidence of hemorrhage and 25% showed edemaxplantation of the probes after 1 day (Schmidt et al., 1993).lthough a large number of the tracks were affected, only 3f the area was actually covered by hemorrhages and euggesting the actual magnitude of the damage to blood vay have been relatively minor. Alternatively, this may heen underestimated by the analytical methods employed

Activated, proliferating microglia appear around the impite as early as 1-day post-implantation (Giordana et al., 1994ujita et al., 1998; Szarowski et al., 2003). Edema and erythroytes remain after 4 days post-implantation, although exuid and cellular debris diminishes after 6–8 days due toction of activated microglia and re-absorption (Stensaas antensaas, 1976). The presence of erythrocyte breakdown pcts (but not hemorrhages) and necrotic tissue can still befter 6 weeks time (Stensaas and Stensaas, 1976; Turner e999). At later time points, some report that typical inflamm

ory cells or hemorrhaging cannot be seen (Yuen and Agnew995), while others have reported observing macrophag

he device brain tissue interface at up to 16 weeks (Stensaand Stensaas, 1976; Turner et al., 1999; Szarowski et al.,iran et al., 2005). As testament to the transitory nature ofechanically induced wound healing response, electrode t

ould not be found in animals after several months whenlectrode was inserted and quickly removed (Yuen and Agnew995; Rousche et al., 2001; Csicsvari et al., 2003; Biran e005), indicating that the persistent presence of the implantents the brain tissue response.

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onths and remained internalized throughout the remaindhe 9-month study (Menei et al., 1994). Immunostaining aroun

needle-like implant revealed scattered reactive microghe initial wound healing response, clustering of microgliaeactive tissue sheath forming around the implant after 2 wnd continued presence of microglia in a tight cellular shea2 weeks post-implantation (Szarowski et al., 2003). Cells notaining for GFAP had adhered to the implant upon its extran the same study.

When macrophages outside the CNS encounter a fobject, they surround it and begin secreting lytic enzyme

ndividual macrophages cannot degrade the object, theseften fuse into multi-nucleated foreign body “giant” cells chcteristic of chronic inflammation. This closely parallelsctivated microglial reaction to electrode implants in the Ceedles made out of a plastic used for tissue mounting (Ara

mplanted into rabbit cortex attracted variably sized mucleated “giant” cells as early as 18 days post-implantStensaas and Stensaas, 1976). These cells were separated frther cortical tissue by a basal lamina and were mostly fdjacent to degraded regions of the plastic needles, sugg

he action of hydrolase activity. A similar layer of tightly coupulti-nucleated giant cells was observed byEdell et al. (1992ith cortically implanted silicon electrodes.A recent study observed persistent ED-1 immunoreac

round silicon microelectrode arrays implanted in rat corteand 4 weeks following implantation that was not observeicroelectrode stab wound controls, indicating that the ph

ype was not the result of the initial mechanical trauma indy probe insertion but was associated with the foreign besponse (Biran et al., 2005). In addition, electrodes explant


at 1, 2, and 4 weeks after implantation were covered with ED-1/OX-42 immunoreactive cells that released MCP-1 and TNF-�in vitro, indicating that inflammation mediated neurotoxic mech-anisms may be occurring at the microelectrode brain tissueinterface.

2.5. Glial scar formation

The most common observation of the long-term CNSresponse to chronically implanted electrodes is the formation ofan encapsulation layer referred to as the “glial scar” (Edell et al.,1992; Turner et al., 1999; Maynard et al., 2000; Shain et al.,2003; Biran et al., 2005). Studies have demonstrated that reac-tive glial tissue surrounds and progressively isolates implantedarrays in a process similar to the fibrotic encapsulation reactionthat is observed with non-degradable implants in soft tissues ofthe body. The development of this encapsulation tissue is limitedto higher vertebrates and has been implicated in the resistanceof the spinal cord and the brain to nerve regeneration after injury(Reier et al., 1983). The purpose of the glial scar remains unclear,but it is thought to play a role in separating damaged neural tis-sue from the rest of the body to maintain the blood–brain barrierand to prevent lymphocyte infiltration (Nathaniel and Nathaniel,1981; Landis, 1994). While soft tissue encapsulation involves avariety of cells and their secreted matrix, reactive astrocytes aret ta 86;E 997T sh trodef singi aka ceb1 ten-s m thr er,1

Turner et al. (1999)used confocal microscopy to show thetime course of astrogliosis. Passive silicon electrodes wereimplanted in the rat cerebral cortex and explanted at 2, 4, 6,and 12-week time points (Fig. 4). At 2 weeks, GFAP stainingrevealed a reactive astrocyte region surrounding the implants thatextended out 500–600�m. This region decreased over time, butthe layer of cells immediately adjacent to the implant becamedenser and more organized suggesting contraction around theimplant. At 2 and 4 weeks, activated astrocytes around theimplant had extended their processes toward the insertion site.The mesh of astrocytic processes became stronger and morecompact at 6 and 12 weeks, as suggested by the fact that removalof the implant did not result in the collapse of cellular processesinto the implantation tract. Both visual and mechanical inspec-tion of the glial sheath suggested that its formation was completeas early as 6 weeks post-implantation and remained intact as longas the implant remained in situ.

A later study by the same group confirmed this time course(Szarowski et al., 2003). This study found a region of diffuseglial activation as imaged by GFAP staining 100–200�m awayfrom the implant site after day 1, and a steady increase of astro-cyte activation to 500�m away from the implant through 1, 2,and 4 weeks. A more compact sheath formed by 6 weeks andremained constant at 12 weeks, with the actual sheath extendingonly 50–100�m around the insertion site. The investigators alsostained for vimentin, which is expressed in reactive astrocytesb timec spa-ta rtantt rodesw xter-n ed tot te thea mayt diese glials new,

F y GF acv proc n.M ial sc

he major component of CNS encapsulation tissue (Schmidt el., 1976, 1997; Schultz and Willey, 1976; Agnew et al., 19dell et al., 1992; Carter and Houk, 1993; McCreery et al., 1urner et al., 1999; Szarowski et al., 2003). Current theorieold that glial encapsulation, i.e. gliosis, insulates the elec

rom nearby neurons, thereby hindering diffusion and increampedance (Schultz and Willey, 1976; Liu et al., 1999; Roitbnd Sykova, 1999; Turner et al., 1999), extends the distanetween the electrode and its nearest target neurons (Liu et al.,999), or creates an inhibitory environment for neurite exion, thus repelling regenerating neural processes away froecording sites (Stichel and Muller, 1998b; Fawcett and Ash999; Bovolenta and Fernaud-Espinosa, 2000).

ig. 4. Time course of glial scar formation at four time points as imaged boid left by the probe extraction before tissue processing. By 6 weeks, theinimal changes between the 6- and 12-week time points indicate the gl

;

e

ut not mature astrocytes. Vimentin expression followed aourse similar to GFAP, but revealed fewer positive cells, aial distribution closer to the implant (25–50�m thick layer), andcompleted sheath at 4 weeks post-implantation. It is impo

o note that in each of the aforementioned studies the electere not functional, that is, they were not connected to an eal electrical connector externalized through and attach

he skull. Such untethered electrodes may underestimactual reactivity caused by electrically active implants that

ransmit forces to the implanted electrode. To date, only stumploying non-specific H&E staining have not observedcar formation (Stensaas and Stensaas, 1976; Yuen and Ag

AP staining. At 2 and 4-week time points, the astrocytic processes fall bk into theesses have interwoven to form a stronger, more dense sheath surroundig the implantar completion within 6 weeks (fromTurner et al., 1999).


1995; Liu et al., 1999), and in a report where both H&E andGFAP were used on the same samples, it was noticed that H&Estaining revealed little gliosis, but GFAP staining clearly showedthe development of a brightly stained astrocytic scar around theimplant (Maynard et al., 2000).

Other factors and cell types may also contribute to the for-mation of the glial scar. Several investigators have reportedthe presence of connective tissue inside this scar similar to theextracellular matrix (ECM) encapsulation seen in wound heal-ing models outside of the CNS (Stensaas and Stensaas, 1976;Liu et al., 1999). Meningeal fibroblasts, which also stain forvimentin, but not for GFAP, may migrate down the electrodeshaft from the brain surface and form the early basis for theglial scar (Cui et al., 2003). Support for this fibroblastic rolein glial scar formation comes from experiments byKim et al.(2004)who compared the cellular response to implants com-pletely surrounded by cortical tissue to transcranial implants thatalso contacted the skull and the meninges (a more accurate modelof functional recording electrode arrays). They found a signifi-cant increase in ECM and connective tissue in the transcranialprobes, as well as a thin layer of GFAP negative/vimentin posi-tive cells surrounding the transcranial probes that was not presentin the implants completely surrounded by brain tissue. Stainingfor ED-1, a microglial marker, confirmed that microglia werepresent within this one to two cell thick layer, and the authorsalso concluded that the presence of ECM suggested meningealfi f thec

2

odes mani welc fromi ed ad trods thei a-ti den-s rouf adnf al tota maga confi wee1 dS

is as lifera thee lacen frome

iridium wire electrodes into feline cerebral cortex and trackedthe stability of single units over time. Histological examina-tion upon explantation revealed that every electrode with stableunit recordings had at least one large neuron near the electrodetip, while every electrode that was not able to record resolv-able action potentials was explanted from a site with no largeneurons nearby. The signal strength declined gradually overtime for failing electrodes, suggesting a gradual remodelingof the environment rather than neuronal death was responsiblefor electrode failure. The study also observed that significantchanges in recording capability occurred in the first 4–8 weekspost-implantation, after which the signals stabilized (such sta-bilization also reported inNicolelis et al., 2003). The authorsof the study speculated that electrode migration through tissueresulted in these changes since active restructuring of the adultmammalian brain is severely limited.

A recent study observed a significant loss of neurons aroundchronically implanted silicon arrays that was not seen in stabwound controls, indicating that the cell loss was associated withthe foreign body response (Biran et al., 2005). Immunostainingrevealed significant reductions of neurofilament and NeuN at theelectrode brain tissue interface that surrounded the implantedelectrodes at 2 and 4 weeks following implantation (Fig. 5). Inthis study although the electrodes were non-functional they weretethered to the skull in a manner similar to working electrodes.In addition, the investigators observed an inverse relationshipb braint m tos vices

2

bilityi eend dif-f re noc fix-a y oft jec-t notu ndo leadst le,i ys,e andi es.A r-m oa ; onet signo s thef nse.P rvedi tedi do o three

broblasts had migrated down the probe from the top oortex.

.6. Neuronal response to implant-induced injury

The density of neurons and their proximity to the electrites are the most accurate barometers of electrode perforn a chronic setting. Unfortunately, this response is not asharacterized as glial scar formation, as it seems to varymplant to implant, and even between electrodes implantifferent sites in the same animal. One explanation of elecignal degradation is the formation of a “kill zone” aroundmplant site resulting from the initial trauma or neuroinflammory events (Edell et al., 1992; Biran et al., 2005). This regions defined by a significantly lower or non-existent neuronality up to some distance away from the electrode. One gound a kill zone of less than 10�m for electrode shafts that hot induced major trauma, with a larger kill zone of 20–60�m

or shafts that experienced minor lateral motion tangentihe brain surface during implantation (Edell et al., 1992). Theuthors suggested a correlation between initial tissue dand kill zone size, although this correlation has not beenrmed by others. Reported kill zone sizes have varied bet�m (no dead zone) and more than 100�m (Stensaas antensaas, 1976; Reier et al., 1983; Turner et al., 1999).An alternative explanation for a loss in signal strength

low regression of neurons away from the electrode. Protion of astrocytes and formation of the glial scar aroundlectrodes could be one mechanism by which gliosis dispeurons near the implant site and pushes cell bodies awaylectrode sites (Edell et al., 1992). Liu et al. (1999)implanted

cel

te

p

e-n

-

s

etween persistent ED-1 staining at the microelectrodeissue interface and loss of neuronal markers, leading thepeculate that persistent activation of microglia at the deurface leads to local neurotoxicity.

.7. In vivo experimental variability

The available evidence suggests there is significant varian electrode performance within experimental groups, betwifferent animal models, investigators, and even between

erent electrodes implanted in the same animal. There aommon handling, packaging, sterilization, implantation,tion, or analytical schemes employed. In addition, man

he experiments are improperly controlled, completely subive and employ too few animals. While such variability isncommon for in vivo work of this kind, the different aften conflicting results obtained from these experiments

o difficulty in drawing meaningful conclusions. For exampn the Edell et al. (1992)study on implanted electrode arralectrode shanks with seemingly identical characteristics

nsertion techniques resulted in significantly different kill zonnother confounding example comes fromRousche and Noann (1998)(Fig. 6), who show an H&E stained image of twdjacent electrode tracks from the same electrode array

rack has healthy neurons growing right up against it and nof an immune response, while the other very clearly show

ormation of a glial scar and a chronic inflammatory respoerhaps the clearest example of this variability was obse

n the in vivo response to plastic “mock electrodes” implann rabbit brain byStensaas and Stensaas (1976)and explantever the course of 2 years. They separated the response int


Fig. 5. Stratification of cellular immunoreactivity using cell-type-specific mark-ers at the microelectrode–brain tissue interface. Representative images collectfrom two adjacent sections of an animal with a 4-week microelectrode implantillustrate the general appearance of the foreign body response characterizedminimally overlapping inflammatory (ED-1) and astrocytic (GFAP) phenotypesadjacent to the implant interface. The area of inflammation and intense astrocytreactivity contains a reduced number of NeuN+ neuronal bodies and a loss ofneurofilament (NF) density. The position of the microelectrode is illustrated bythe orange oval (drawn to scale) at the left of each image. Images were capturein grayscale and pseudocolored for illustration (fromBiran et al., 2005).

Fig. 6. An example of in vivo variability commonly encountered in implantliterature. Arrows show adjacent tracks from two electrodes on the same array.The track on the right clearly shows heavy matrix deposition while that on theleft seems to have no tissue reaction (fromRousche and Normann, 1998).

types: Type 1 was characterized by little to no gliosis with neu-rons adjacent to the implant, Type 2 had a reactive astrocyte zone,and Type 3 exhibited a layer of connective tissue between thereactive astrocyte layer and the implant, with neurons pushedmore than 100�m away. All three responses are well docu-mented in the literature; however this study found that the modelelectrodes produced all three types of reactions simultaneously,depending on where along the electrode one looked. Althoughthese studies clearly suffer from the insensitivity of the chosenhistopathological approach (non-specific tissue staining), it isclear that only a broader view of the literature can yield mean-ingful conclusions in the face of such experimental variance.

3. Current electrode implant systems

The problems inherent in chronic recording electrode designhave precluded the development of a “gold standard” electrodeagainst which testing is performed. Historically, neurobiologyresearch has used single wire or glass micropipette electrodes torecord individual neuron waveforms in acute experiments. How-ever, the need to access populations of neurons and the desire ofresearchers to monitor neuron networks over time has added anew focus on arrays of wires, silicon shafts and other more com-plex micromachined silicon recording systems capable of highdensity sampling. Chronic implantation has also generated var-i e andt alientf s theyr

3

inge andw e ofa ,11 d

ed

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d

ous surgical techniques aimed at reducing electrode failurhe foreign body response. This section details the most seatures of electrode array design and surgical techniques aelate to biocompatibility.

.1. Multiple electrode types

Of the two main types of electrode arrays currently bexplored, microwire electrodes have the longest historyidest use in the field. Microwire electrodes are wires madconducting metal, such as platinum, gold (Yuen and Agnew

995), tungsten (Williams et al., 1999b), iridium (Liu et al.,999), or stainless steel (Nicolelis et al., 2003), that are coate


Fig. 7. (a) Wire electrode arrays implanted in macaque monkey cortex. (b) Layout of six such wire arrays in macaque monkey cortex (fromNicolelis et al., 2003).

with a non-cytotoxic insulator material. The tip of the wire is notinsulated and can receive electronic signals from the surroundingneurons. In an effort to better separate single unit from multi-unit activity, experimenters often record from two (stereotrodeconfiguration;Mcnaughton et al., 1983) or four (tetrode config-uration;Gray et al., 1995) closely spaced microwires to allowrelative signal strength to act as another parameter in single unitidentification. Finally, microwires can be arranged in arrays toaccess the large numbers of neurons necessary for neuropros-thesis control. The number of wires used in a single implantedarray has ranged from 4 (Yuen and Agnew, 1995) to over 100(Nicolelis et al., 2003). A clear advantage of using microwireelectrodes is the ease in array fabrication compared to moresophisticated silicon arrays, which are discussed later. Althoughmicrowire arrays are simpler, their performance in recordinghigh numbers of single units often exceeds the quality of record-ings obtained from silicon-based electrodes.Nicolelis et al.(2003)chronically implanted 10 microwire arrays into macaquemonkey cortex for a total of 704 microwires, and were able torecord 247 individual cortical neurons in a single session from384 of the microwires (Fig. 7). While the number of units var-ied from day to day and between monkeys, the yield of unitsper recording site was much greater than that of the typical sil-icon array. Furthermore, microwire arrays can access deeperbrain structures, but the precise location of the electrode tipsand the interelectrode spacing cannot be controlled as the non-h ringi

cond , thn domn fora giest ovee inint elect meoe ipke

et al., 2003). However, despite substantial technologicaladvances in their design, many such devices are unreliable forchronic recording applications in the CNS (Liu et al., 1999).

Silicon photolithographic processing allows for unsurpassedcontrol over electrode size, shape, texture, and spacing, allow-ing multiple recording sites to be placed at variable heights on asingle electrode shank. Such control provides the experimenterwith absolute knowledge of the recording location, the ability toplace the recording sites at different depths to suit the geometryof the neural system under study, and a larger overall numberof recording sites on a smaller volume than is possible on wirearrays or bundles (Kewley et al., 1997). Circuits can be integrateddirectly on the probe for better signal acquisition, and on-chipmicroelectromechanical systems (MEMS) add additional possi-bilities, such as heating elements and microfluidics (Chen et al.,1997; Bai and Wise, 2001). Further decreases in electrode sizesand increases in recording site densities are currently limited byconnectors and on-board systems that are unable to handle thehundreds of possible leads on a single array (Campbell et al.,1991; Maynard et al., 2000; Bai and Wise, 2001; Kipke et al.,2003).

The designs of silicon-based electrode arrays vary betweeninvestigators and research manufacturing centers (Fig. 8) (Edellet al., 1992; Turner et al., 1999; Bai and Wise, 2001; Csicsvariet al., 2003; Szarowski et al., 2003). Nevertheless, two partic-ular silicon electrode array designs have attained prominencei byN years( narde onw s in at rdingt riouss from1 to2 es;N

iver-s ogy

omogeneous nature of brain tissue will bend microwires dumplantation (Edell et al., 1992).

Although most neuroscience research continues to beucted using these well-established microwire electrodesext generation of electrode arrays being developed is preantly silicon based. Silicon micromachined electrodes allowmore complex design and thus greater flexibility in strate

o minimize the foreign body response and greater controllectrode placement. The emergence of silicon micro-mach

echnology has yielded increasingly smaller and higherrode count arrays capable of recording from greater voluf neural tissue with improved spatial discrimination (Draket al., 1988; Branner et al., 2001; Csicsvari et al., 2003; K

-ei-

rg-s

n the field. The Utah Electrode Array (UEA) developedormann and co-workers has been in use for over 15

Campbell et al., 1991; Rousche and Normann, 1998; Mayt al., 2000). The UEA is created from a single block of silichich, through etching, doping, and heat treatment, result

hree-dimensional array of needle-like electrodes with recoips. It has been made in 25 and 100 shank versions of vahapes, with each shank 1.5 mm in length and ranging00�m at its base to less than 1�m at the tip (as compared5–50�m diameter and up to 8 mm in length for microwiricolelis et al., 2003).The other prominent electrode design comes from the Un

ity of Michigan Center for Neural Communication Technol


Fig. 8. Various designs of silicon micromachined electrode arrays. (a) Arrow points to a “well” included in the electrode design for bioactive molecule incorporation.Multiple electrode sites are present on each electrode shank (fromKipke et al., 2003). (b) Utah Electrode Array formed from a single block of silicon (fromRouscheand Normann, 1998). (c and d) Multiple planar arrays of “Michigan” electrodes are stacked together to create a three-dimensional array (fromBai and Wise, 2001).

(Cui et al., 2001, 2003; Kipke et al., 2003; Szarowski et al.,2003). Unlike the UAE, the Michigan probes are planar arraysof electrode shanks made from a single thin sheet of silicon,but each 2 mm long, 100�m× 15�m shank has several record-ing sites placed at user-defined locations along the shank. Themanufacturing process is similar to that used to create multi-layer silicon microchips and requires eight masks, but results inhundreds of user-defined, reliable, and regular electrodes froma single block of silicon. A different group has developed a sim-ilar design with four recording sites per shank using ceramicrather than silicon for their electrodes, resulting in mechani-cally stronger electrodes with comparable recording capabilities(Moxon et al., 2004a,b).

3.2. Materials used for the insulating layer

Both microwire and silicon-based electrode array sys-tems require an insulation layer to shield the electrodesfrom unwanted electrical signals. The vast majority oftissue–electrode contact is with the insulating layer, so this mate-rial must be non-toxic and should act to reduce the foreignbody response. Several different materials identified as mini-mally toxic have been used to coat electrodes. A simple coatingof Teflon or S-isonel, a high temperature polyester enamel sim-ilar to Teflon, has been used with great success to coat wire

electrodes (Kennedy, 1989; Nicolelis et al., 2003). Resins, suchas Epoxylite, have also been used successfully (Liu et al., 1999).Plasma-deposited diamond-like carbon (DLC) has recently beendemonstrated in vitro as both a chemically inert insulator and asa good substrate for biological molecule attachment to controlthe foreign body response, although it has not been tested in vivo(Ignatius et al., 1998; Singh et al., 2003).

In addition to the normal silicon nitride or silicon dioxideinsulation deposited during the fabrication of micromachinedsilicon electrodes (Campbell et al., 1991; Kipke et al., 2003),polyesterimide, or more commonly, polyimide is often used tocoat silicon-based implants (Campbell et al., 1991; Yuen andAgnew, 1995; Williams et al., 1999b; Bai and Wise, 2001). Leeet al. (2004)reported that fibroblasts spread, adhered, and grewon polyimide electrode surfaces with no difference from tissueculture polystyrene controls. The electrodes were also mountedon a thin (5–10�m) silicon substrate to aid in electrode insertionthrough the pial membrane of the brain (Lee et al., 2004). Theflexibility of polyimide may improve the mechanical impedancemismatch between a rigid electrode and soft tissue resulting intissue damage if micromotion of the electrode occurs (Rouscheet al., 2001; Kim et al., 2004). Rousche et al. (2001)eliminatedsilicon completely and created a flexible electrode out of poly-imide and gold, where the gold recording sites and leads weresandwiched between two layers of polyimide. A major drawback


to this design was that the electrodes were not stiff enough topierce brain tissue on their own, so implant sites had to be createdwith wire or a scalpel before insertion.

3.3. Electrode insertion and implantation procedure

Many studies have attributed biologically induced electrodefailure to the initial trauma of implantation, leading to a varietyof strategies to minimize this early trauma in the hope of lim-iting the subsequent complications. Unfortunately, since eachgroup of investigators works with a different electrode system,a different animal model, and a different set of hands, thereis very little consensus regarding the optimal way to implantchronic recording electrodes and a shortage of well-controlledquantitative studies. Approaches differ on the speed of electrodeinsertion, the method of insertion, the importance of limitingmicromotion, and the depth of insertion.

Experimenters use a wide range of insertion speeds for elec-trode implantation. One theory holds that slow insertion allowsfor neuronal tissue to adjust to the implant, thus minimizingthe damage caused by the electrode. In the experiments con-ducted byNicolelis et al. (2003), a 100�m/s microwire electrodeinsertion rate was cited as a major factor behind the unusuallylarge number of single units recorded in the study. However,other groups have reported problems with slow insertion, suchas catching of the tissue and dural dimpling (Edell et al., 1992).T izest ue it ta loc-i andm l.,1 edsi ;S

ringo sa 999Wm stomb hindt riee eent ial bc usea evern poinK aftr g top dy.

e fore ttachm linkb idec cabw

another group used a flexible printed circuit board to overcomethe perceived problem of micromotion (Kipke et al., 2003).Another set of experiments revealed that adhesions had formedbetween a UEA implant and the dural tissue above the brain,possibly due to a robust foreign body response mediated bymicromotion of the implant relative to the dura (Rousche andNormann, 1998). Placing a Teflon sheet between the array andthe dura, and a sheet of Gore-Tex® between the dura and thecranium in subsequent experiments significantly improved theperformance of the electrode array over the course of 9 months(Maynard et al., 2000). The Teflon sheet also affected implantmigration within cortical tissue, a cause of signal degradationcited by another group that had observed longer microwires los-ing neuronal signals before their shorter counterparts within thesame electrode array (Liu et al., 1999).

4. Strategies to minimize the immune response toimplanted electrodes

With different electrode array technologies, machiningoptions, biocompatible materials, and implantation proceduresavailable, various groups have altered the design of electrodes inan attempt to minimize or evade the immune response. Investiga-tors better acquainted with the molecular biology of the neuralenvironment have also added bioactive agents to the materialscience repertoire of electrode designers. This intersection ofn rablep

4

gingm ewed( 00T ingt NSa nicala a haveb onse.S lant( ofeHp nd tipg iliconi nser-t -1i s oft ec o sur-f harpb ods( ningr s (oft ring,a all oft luded

he other school of thought is that a rapid insertion minimrauma, since the force of the insertion cuts through the tisshe array’s path, but does not affect nearby tissue (Campbell el., 1991). Groups using the UEA have found that a high ve

ty approach (8.3 m/s) prevents cortical surface dimplinginimizes tissue damage (Campbell et al., 1991; Schmidt et a993; Maynard et al., 2000). Other groups use insertion spe

n between (i.e. 2 mm/s) these two extremes (Turner et al., 1999zarowski et al., 2003).The method of insertion has also been a source of diffe

pinions. Some groups insert the electrodes by hand (Stensaand Stensaas, 1976; Yuen and Agnew, 1995; Liu et al., 1illiams et al., 1999b; Szarowski et al., 2003), while other utilizeicrodrives to make the delicate insertions and cite such cu

uilt or commercially available devices as a major factor behe success of their experiments (Maynard et al., 2000; Csicsvat al., 2003; Nicolelis et al., 2003; Szarowski et al., 2003). Edellt al. (1992)suggested that maintaining an alignment betw

he electrode shaft and the axis of insertion was essentalculating that a 1◦ misalignment in a 1 mm insertion can ca17�m slash through the tissue at the insertion site, how

o experimental evidence was offered to substantiate thisewley et al. (1997)point to a misalignment of the probe sh

elative to the insertion axis as the major factor contributinoor signal strengths of some implanted probes in their stu

Several strategies have been employed to minimize thign body response via adjustments to electrode implant aents. One source of electrode micromotion can be theetween the electrode and its connector. A flexible polyimonnection between the rigid electrode and the connectoras utilized in one set of experiments (Lee et al., 2004), while

n

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le

eural immunobiology and electrode design holds consideromise for developing reliable and useful probes.

.1. Material science strategies

The materials science and biocompatibility of packaaterials for sensors outside of the CNS has been revi

Sharkawy et al., 1997, 1998a,b; Wisniewski et al., 20).he majority of previously attempted strategies for limit

he immune reaction to electrodes implanted within the Clso revolve around material science and physical/mechapproaches. Electrode size, shape and cross-sectional areeen modified to elicit the smallest possible tissue respome reports give significant weight to the texture of an imp

Rousche et al., 2001). Other studies stress the importancelectrode tip shape (Edell et al., 1992; Nicolelis et al., 2003).owever, a recent study bySzarowski et al. (2003)down-layed the importance of electrode shape, size, texture, aeometry. The study compared the immune response to s

mplants of different sizes, surface characteristics, and iion techniques (Fig. 9) through GFAP, vimentin, and EDmmunostaining over the course of 12 weeks. Electrodehree sizes (2500, 10,000, and 16,900�m2 cross-sections), threross-sectional shapes (trapezoid, square, and ellipse), twace textures (smooth and rough), two tip geometries (slade-like point and rounded tip), and two insertion methhand and precision drive) were tested in rats. Glial staievealed that while there were minor temporal differencehe order of 1–3 weeks) in the time course of the glial scart 6 and 12 weeks post-implantation the tissue response to

hese electrodes was essentially identical. The study conc


Fig. 9. Different sizes, shapes, and cross-sections of electrodes that produced the same foreign body response and glial scar, suggesting that strategies to improvebiocompatibility through purely structural changes in the implant may be ineffective by themselves (fromSzarowski et al., 2003).

that while the various geometries may affect the initial woundhealing response, glial scar formation was not affected, how-ever, the lack of observable differences may have been due tothe low animal number, a lack of controls, and variability ofresponse form animal to animal. Although different materialswere not tested by Szarowski et al., other experiments have notshown any significant reduction in the immune response withvarious metals (Ignatius et al., 1998) or other materials, such asDLC (Singh et al., 2003). Such studies support the shift froma materials science strategy in evading the immune response tostrategies focusing on the molecular and cellular biology of theimmune response.

4.2. Bioactive molecule strategies

With material science strategies failing to eliminate glialencapsulation, a failure that parallels sensor implants outsidethe CNS, a number of investigators are examining approachesthat manipulate the biological response. Since the proximity ofneurons correlates directly with signal strength, strategies thatattract, attach, or preserve neurons near recording sites couldminimize the effect of the immune response on electrode perfor-mance. Such strategies have focused on coating electrodes withbioactive molecules, such as cell adhesion promoting polypep-tides or proteins. Neurons use environmental cues to grow,migrate, and stay viable. Some of these cues are in the form ofp em-b tactce pept -ca -D e-s thesp et a

found that poly-d-lysine (a synthetic polypeptide that enhancesneural-cell adhesion) and laminin, when co-absorbed on variousmetals and glass, greatly improved cell attachment, spreading,and growth as compared to uncoated metals. Polyimide, the com-mon insulating material discussed earlier, is also amenable tosurface modification with bioactive molecules. Recent studies byMartin have focused on using conducting polymers of polypyr-role and poly(3,4-ethylenedioxythiophene) to provide bettercontact and a larger surface between the electrode and adjacentneuronal tissue (Cui et al., 2001, 2003; Yang and Martin, 2004).These polymers can be “grown” in a controlled manner throughan electrochemical process at the electrode recording sites andcan easily incorporate bioactive molecules. Electron micro-graphs of the sites reveal finger-like fibers of polymer growingout of the gold electrodes, creating a “fuzzy” surface with alarge surface area to maximize neuron–electrode interactions.Cui et al. (2001, 2003)incorporated the YIGSR peptide frag-ment from laminin into the polypyrrole coated recording sites.The peptide-coated sites supported more neuron attachment invitro compared to peptide-free sites (Cui et al., 2001). When thepeptide-incorporated and peptide-free electrodes were chron-ically implanted in guinea pigs, 83% of peptide-incorporatedversus 10% of peptide-free electrode sites showed evidence ofneuronal process proximity after 1 week, but the two types ofimplants produced similar recordings and exhibited similar glialscar reactions (Cui et al., 2003).

ed toe owo xtenda of themt hort n, oree fer-e ed to

olypeptide motifs on extracellular matrix proteins or on mrane bound molecules of neighboring cells. In addition to inell adhesion proteins, such as collagen and fibronectin (Ignatiust al., 1998), several groups have employed cell adhesion

ides, such as RGD (Kam et al., 2002), found in many extraellular proteins, YIGSR (Kam et al., 2002; Cui et al., 2003),nd IKVAV (Kam et al., 2002), found in laminin and KHIFSDSSE (Kam et al., 2002), found on NCAM (neural-cell adhion molecule). Studies to establish neuronal reactions toroteins and peptides have been conducted in vitro. Ignatius

-

el.

Bioactive molecule surface coatings have also been usither attract or repel glial cells. Neurons in culture will grn astrocyte monolayers and neuronal processes will elong tracks provided by astrocytes, oftentimes regardlessaterial beneath the astrocytes (Biran et al., 1999, 2003). Attrac-

ion of astrocytes and other glial cells could potentially anche electrodes in the neural tissue and prevent micromotioven block negative components of the glial response (Kamt al., 2002). Kam et al. found that astrocytes adhere prentially to glass substrates covered with NCAM as compar


other adhesion molecules. A separate study found that astrocytesalso prefer silk-like polymer fragments with fibronectin domainsover laminin YIGSR domains, which are preferred by neurons(Cui et al., 2001). Robust growth of glial cells on laminin, colla-gen, and fibronectin-coated surfaces was also observed (Ignatiuset al., 1998). Shain and co-workers have attempted to control andpattern astrocyte adhesion through the deposition of hydropho-bic and hydrophilic self assembling monolayers of organosilanesusing photolithography and microcontact printing (St. Johnet al., 1997; Craighead et al., 1998; Kam et al., 1999). Pho-tolithography also allows the manufacture of silicon pillar arraysof varying dimensions (Craighead et al., 1998; Maynard et al.,2000; Turner et al., 2000), which are consistently preferred tosmooth silicon by glial cells in vitro. Other studies have triedto prevent astrocyte adhesion in an effort to reduce or eliminateglial scar formation.Singh et al. (2003)found that a dextrancoating of DLC-poly-lysine surfaces reduced glial cell adhe-sion more than 50-fold to 1.41± 1.23% of control surfaces.The study however, was performed using cell lines in vitro, anddid not address how such a coating would affect neurons andmicroglia.

Whether such approaches work in vivo has not been deter-mined. The results indicate a promising direction for research,but it will be necessary to find bioactive molecules that main-tain neurons in proximity to the recording surface, minimizeastrogliosis, and eliminate chronic microglial activation. To date,I ;K ,1 ainc

rdingp actoo owtht ophf n ofn eralbOy ctrow upoc uralp de tia . Thr wirea nthsw wedn wtha con-t vely,t ttrat esigs frome calee ntinp rooo lublef ls.

Standard wound healing suppression and immunosuppres-sion techniques are also options for minimizing the initialimmune response and perhaps even the glial scar formation. Bothlocal and systemic administration of corticosteroids and otherdrugs has been shown to reduce the wound healing response forimplants outside of the CNS (Hickey et al., 2002; Yoon et al.,2003). Upon implantation of the UAE into cats byMaynard et al.(2000), two doses of Dexamethasone, a potent suppressant of thewound healing response, were administered 12 h before surgeryand again during surgery to control cortical edema, resultingin considerable improvements in implant performance. Whilepromising, the success was likely due to the implantation of thepreviously mentioned Teflon sheet.Shain et al. (2003)foundthat peripheral injections of Dexamethasone at the time of elec-trode insertion greatly attenuated glial scar formation at 1 and 6weeks as shown by GFAP staining. Some attenuation was seenwith local release of Dexamethasone from implanted poly(ethyl-vinyl) acetate strips, but at 6 weeks post-implantation the effectwas minor. A peripheral injection of Cyclosporin-A, anotherpotent anti-inflammatory agent, in the same study seemed toincrease the glial response.

With the surprising success of Kennedy’s cone electrodecame an equally surprising failure. When the same micropipetteelectrode was filled with a solution of neural growth factor(NGF) in various concentrations to mimic the effect of thesciatic nerve used in the successful trials, no ingrowth waso da utedt allyo or-t t canc oat-t theC ac-t liv-e andpWe pt,b ro-g rugd est rodes 0c ated2 pa-r horss ,1

5b

cadeo bei log-i arated

KVAV and YIGSR polypeptide fragments (Cui et al., 2001am et al., 2002) and poly-d-lysine on metal (Ignatius et al.998) have elicited different behavior from the various brell types, but a “magic bullet” has not been found.

Aside from helping neurons adhere to the electrode recoads, a more active approach may be to release growth fr chemoattractants to promote neuronal survival and gr

owards the electrode. Numerous chemoattractants and tractors act during brain development to facilitate the creatioeural circuits, but growth factor based strategies have geneen disappointing in the adult CNS (Stichel and Muller, 1998a).ne promising result was obtained byKennedy (1989)over 15

ears ago when he seeded a standard glass pipette eleith a piece of sciatic nerve and observed what happenedhronic implantation of the so-called “cone” electrode. Nerocesses from surrounding neurons grew into the electrond this ingrowth could be tracked by the recorded signalsecorded SNR was often 5–10 times that obtained withnd silicon electrode arrays and continued for over 12 mohile control electrodes without the sciatic nerve insert shoo neurite ingrowth. Furthermore, the amount of this ingrond the resultant single versus multi-unit activity could be

rolled by adjusting the size of the cone tip opening. Putatihe sciatic nerve tissue released growth factors or chemoaants that caused neurite ingrowth while the electrode dhielded the neurites from the foreign body response andxternal electrical noise. Clearly this design cannot be sasily to electrode arrays for neural prosthetics, and explaieces of PNS tissue for such an array is prohibitive. As a pf concept however, Kennedy showed the promise of so

actors in attracting and retaining high quality neural signa

rs

ic

ly

den

pe

,

c-n

dgf

bserved (Kennedy, 1989). Instead, a cystic cavity formeround the glass cone electrode, which the author attrib

o hyperplastic growth of surrounding tissue that eventuutgrew its blood supply. This failure highlights the imp

ance of developing appropriate drug delivery systems thaapitalize on the positive effects of growth factors or chemractants. The normal difficulties of delivering drugs toNS are compounded with the difficulties in delivering bio

ive molecules over long implantation periods. Current dery methods include polypyrrole grown on electrode padsolypeptide-doped polyimide coatings (Cui et al., 2001, 2003).ells etched into the polyimide electrode developed byRousche

t al. (2001)were filled with dextran as a proof of conceut could potentially hold other diffusible compounds or hydels. Chen et al. (1997)have attempted to address this delivery problem with the development of “puffer” prob

hat incorporate microfluidic channels inside of the electhank. These bulk machined silicon probes have multiple 1�mhannels for chemical and drug delivery from orifices situ.5�m from recording sites. There is currently no gating apatus at the orifices to control of fluid release, but the autuggest a shutter mechanism is being explored (Chen et al.997).

. Perspectives on the current state of the electrodeiocompatibility field

From the large amount of data collected over the past den intracortical implant biocompatibility, several trends can

dentified for designing future experiments. The complex biocal reaction against the implanted electrodes can be sep


Fig. 10. Possible mechanisms and time course of cellular response to an implant. Neurons are in pink, astrocytes in red, microglia in blue, and vasculature in purple(from Szarowski et al., 2003).

into two immune responses (Fig. 10). The acute phase is a 1–3-week long process in which microglia play a dominant role inresponse to the insertion trauma. It is unclear how the intensityof the acute response affects subsequent events, which involveboth reactive astrocytes and chronically activated microglia.The astrocytic response begins at the time of insertion and isgenerally completed by 6–8 weeks post-implantation with thedevelopment of an encapsulating glial scar. Neuron viabilityclearly decreases following device insertion, but the questionremains whether the neurons that survive the acute reaction andremain in proximity of the chronic foreign body response remainelectrically active or viable in the presence of persistent inflam-mation. The astrocytic scar remodels nearby tissue, thus furtherseparating neurons from the recording electrodes, and possiblyincreasing electrode impedance. All of these factors most likelycontribute to inconsistent performance of recording electrodesand eventually result in a loss of recorded extracellular poten-tials. The complexity of this response, coupled with a lack ofwell-controlled experiments and in vitro model systems of reac-tive gliosis hinders the development of biointeractive strategies.Without a better understanding of the roles that cells, solublefactors, and the extracellular matrix play in both the acute andchronic responses, strategies, such as systemic immunosuppres-sant application and the incorporation of PNS explants, willremain promising but not clinically applicable. One importantstep will be developing in vitro cell culture models of reactiveg rate-g

pro-c bati blea astc r anm thisw sur-f dM supp tantD polyi notb

Until recently, attempts to improve electrode reliability inchronic implants have centered on reducing insertion damageand limiting micromotion in vivo (Edell et al., 1992; Yuenand Agnew, 1995; Maynard et al., 2000; Rousche et al., 2001;Shain et al., 2003; Lee et al., 2004). While this approach mayhave merit, more well-controlled experiments are needed tofirmly establish how micromotion contributes to the problemor how it relates to different designs. Alternatively, concen-trating efforts to eliminate the chronic reaction rather thantrying to minimize the acute reaction may meet with greatersuccess (Reier et al., 1983; Turner et al., 1999; Shain et al.,2003).

The Szarowski et al. (2003)study of the effect of shape,size, and texture on the immune response suggested that a non-biological approach to electrode design may not be sufficientto overcome the biological hurdles of chronic electrode implan-tation. Consequently, an electrode design that does not couplepharmacological delivery of bioactive molecules or utilize site-specific drug release systems may not overcome the body’simmune response. These notions are likely responsible for theshift in next generation electrode array designs from microwireelectrodes to silicon-based arrays. However, until these siliconarrays prove successful in vivo the neuroscience field will con-tinue to use the simpler, more constrictive, but more reliable,microwire arrays (Nicolelis et al., 2003).

Finally, it is important to consider the duration over whichm onicb notd le inm t pastt hasb waysb ivoe lantr y aref iesi en-d withn elec-t Yuena 003;

liosis to facilitate a more rational design of therapeutic sties.

Once a better understanding of the underlying biologicalesses is obtained, several strategies may be used to commmune response. Some researchers believe it is possittract neuronal processes to the electrode sites before theytic scar develops, enveloping the neurites within the scaaintaining a signal. Martin and co-workers are pursuingith conductive polymer/biomolecule blends grown on the

ace of electrode recording pads (Cui et al., 2001, 2003; Cui anartin, 2003). At the same time neurons are being attracted,ression of the astrocyte scar formation may prove imporextrans or other non-adhesive molecules embedded in

mide coatings have the potential to fulfill this role, but haveeen tested in vivo (Singh et al., 2003).

thetoro-d

-.-

any studies are conducted. Clinical applications of chrrain implants must remain active for at least a year ifecades. While decade-long experiments are not feasibost research environments, studies should be taken ou

he 12-week mark to ensure that a full immune responseeen mounted against the implant. In vitro studies have aleen helpful in guiding the direction of more intensive in vxperiments, but in the long time periods involved in impejection, cell culture experiments are of little use unless theollowed by implantation work. Benchmarking in such studs also critical to accurately control for experimenter depent factors. Many biocompatibility studies are conductedon-functional electrodes rather than with active recording

rodes (Stensaas and Stensaas, 1976; Edell et al., 1992;nd Agnew, 1995; Turner et al., 1999; Szarowski et al., 2


Biran et al., 2005) that may more accurately model the braintissue response. As such, it is difficult to evaluate electrodeperformance relative to established electrode arrays, and futurestudies need to benchmark against an established electricallyactive design to augment the comparison studies already inthe literature (Rousche and Normann, 1998; Liu et al., 1999;Maynard et al., 2000; Kipke et al., 2003).

In addition to teaching us about the normal operation of thebrain, evidence is accumulating that intracortical multi-channelrecording interfaces have significant potential to provide controlsignals for neuroprosthetic devices ranging from motor controlin paralyzed patients to restoring sensory function in the auditoryand visual systems (Kennedy and Bakay, 1998; Donoghue, 2002;Serruya et al., 2002; Taylor et al., 2002; Vaughan et al., 2003).The enormous benefits that these implants could deliver haveimpelled the scientific community to focus on the clinical prob-lems involved in long-term implantation. This focus has paiddividends in the form of more reliable electrode arrays, moreaccurate measurement techniques, and a better understanding ofthe processes involved in implant rejection in the CNS. Realizingthat they must leverage each others’ strengths to create a reli-able implant system, neuroscientists specializing in recordingneuronal signals and engineers designing implantable electrodesystems are increasingly collaborating with each other at var-ious stages of experimentation and design. Such collaborationhas yielded a general view that strategies to prevent the foreignb arta min-i ewf s cad car-r NSi 99;B 200S ta odyr willa lesi

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Bai Q, Wise KD. Single-unit neural recording with active microelectrodearrays. IEEE Trans Biomed Eng 2001;48:911–20.

Banati RB, Gehrmann J, Schubert P, Kreutzberg GW. Cytotoxicity ofmicroglia. Glia 1993;7:111–8.

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Biran R, Martin DC, Tresco PA. Neuronal cell loss accompanies the braintissue response to chronically implanted silicon microelectrode arrays.Exp Neurol 2005;195:115–26.

Biran R, Noble MD, Tresco PA. Characterization of cortical astrocyteson materials of differing surface chemistry. J Biomed Mater Res1999;46:150–9.

Biran R, Noble MD, Tresco PA. Directed nerve outgrowth is enhanced byengineered glial substrates. Exp Neurol 2003;184:141–52.

Bovolenta P, Fernaud-Espinosa I. Nervous system proteoglycans as modula-tors of neurite outgrowth. Prog Neurobiol 2000;61:113–32.

Branner A, Stein RB, Normann RA. Selective stimulation of cat sciaticnerve using an array of varying-length microelectrodes. J Neurophysiol2001;85:1585–94.

Campbell PK, Jones KE, Huber RJ, Horch KW, Normann RA, Silicon-BasedA. 3-Dimensional neural interface—manufacturing processes for an intra-cortical electrode array. IEEE Trans Biomed Eng 1991;38:758–68.

Carmena JM, Lebedev MA, Crist RE, O’Doherty JE, Santucci DM, Dimitrovand

C sing

C isibi-

C e forEng

C M, etntral.

C t al.on-

C al.poly-

C andSens

C of–87.

D n inter-

D ce ofntra-

E within

E m-rans

E gesn and

ody response and glial scar formation are a necessary pny implant design. While earlier strategies have dealt with

mizing and varying the implant footprint within the CNS, a nocus on bioactive strategies is emerging. This research focuraw upon the decades of work in the prevention of glial sing and the initiation of neuronal regeneration in traumatic Cnjuries (Stichel and Muller, 1998b; Fawcett and Asher, 19ovolenta and Fernaud-Espinosa, 2000; Hermanns et al.,chmidt and Leach, 2003; Yick et al., 2003). It is our hope tharational approach to limiting or eliminating the foreign b

esponse in the CNS through biologically active mediatorsugment current implant designs and overcome the obstac

ntroducing chronic CNS implants into the clinic.

cknowledgements

The authors are grateful for stimulating and informaiscussions with M.L. Block and J.S. Hong of NIEHS. Tork was supported by a National Science Foundation Gate Research Fellowship (V.S.P.). Support from the Nat

nstitutes of Health is also gratefully acknowledged (W.Mnd P.A.T.).

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response of brain tissue to chronically implanted neural electrodes

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