Implantable Bio Electronic Interfaces for Lost Nerve Functions

8/8/2019 Implantable Bio Electronic Interfaces for Lost Nerve Functions

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IMPLANTABLE BIOELECTRONIC INTERFACES FOR LOST

NERVE FUNCTIONS

P. HEIDUSCHKA* and S. THANOS*6

*University Eye Hospital Munster, Experimental Ophthalmology, Domagkstrae 15,D-48149 Munster, Germany

(Received 6 January 1998)

Abstract Neuronal cells are unique within the organism. In addition to forming long-distance connec-tions with other nerve cells and non-neuronal targets, they lose the ability to regenerate their neuritesand to divide during maturation. Consequently, external violations like trauma or disease frequently

lead to their disappearance and replacement by non-neuronal, and thus not properly functioning cells.The advent of microtechnology and construction of articial implants prompted to create particulardevices for specialised regions of the nervous system, in order to compensate for the loss of function.

The scope of the present work is to review the current devices in connection with their applicabilityand functional perspectives. (1) Successful implants like the cochlea implant and peripherally implanta-ble stimulators are discussed. (2) Less developed and not yet applicable devices like retinal or corticalimplants are introduced, with particular emphasis given to the reasons for their failure to replace verycomplex functions like vision. (3) Material research is presented both from the technological aspect andfrom their biocompatibility as prerequisite of any implantation. (4) Finally, basic studies are presented,which deal with methods of shaping the implants, procedures of testing biocompatibility and modi-cations of improving the interfaces between a technical device and the biological environment. Thereview ends by pointing to future perspectives in neuroimplantation and restoration of interrupted neur-onal pathways.# 1998 Elsevier Science Ltd. All rights reserved

CONTENTS

1. Introduction 4342. Implantable electrodes 435

2.1. General remarks 4352.2. Electrodes implanted without nerve cut 435

2.2.1. Cu electrodes 4352.2.2. Penetrating electrodes 437

2.3. Electrodes for regenerating nerves 4383. Biocompatibility 441

3.1. General remarks 4413.2. Implant materials 4413.3. Response to implantation 4423.4. Surface modication 443

3.4.1. Surface roughening 4433.4.2. Chemical modication 4433.4.3. Specic modications 444

4. Auditory implants 4454.1. Brain stem implants 4464.2. Cochlea implants 446

5. Visual implants 4495.1. Retinal prostheses 4495.2. Cortical prostheses 451

6. Other neuroprosthetic implants 4526.1. Bladder stimulation 4526.2. Stimulation of spinal cord and brain 454

7. Final conclusions 455References 455

Progress in Neurobiology Vol. 55, pp. 433 to 461, 1998# 1998 Elsevier Science Ltd. All rights reserved

Printed in Great Britain0301-0082/98/$19.00

PII: S0301-0082(98)00013-6

6Author for correspondence. Tel.: 49 251 83 56915; Fax: 49 251 83 56916; e-mail: [email protected].

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ABBREVIATIONS

ABI Auditory brain stem implantANN Articial neural network

BDNF Brain-derived neurotrophic factorCNS Central nervous systemECM Extracellular matrixFES Functional electrical stimulationIFN InterferonIL Interleukin

MEA Microelectrode arrayNT-3 Neurotrophin-3

PET Polyethylene terephthalatePTFE PolytetrauoroethyleneRGC Retinal ganglion cellsRP Retinis pigmentosaTNF Tumour necrosis factorVEP Visually evoked potential.

1. INTRODUCTION

Diseases and accidents associated with damage ofnerves, in particular within the CNS, often have dra-matic consequences. The rst reason is that appro-priate regeneration of the central neuronalconnections and restoration of synaptic connectionsare not possible in most cases. This results in failureof functional recovery. The second reason is thatdying and disposed neurons cannot be replaced bynew neurons. This cascade of interactive eventsresults in glial proliferation and in inadequaterepair, called gliosis. In contrast to CNS, bundles ofthe peripheral nerve system like in the limbs displaydierent responses to experimental or accidental cut:the axons can regenerate, the cell bodies do notdegenerate and there is virtually no need for gliosisand replacement of the lost neurons. Decits alongperipheral nerves may be reconstructed spon-taneously or surgically, and physiological propertiesof the function are regained after regeneration of

axons and synapse formation. The nal functionaluse of the corresponding target deserves some train-ing, as directly associated with synaptic reorganisa-tion, potentiation and trophic inuences at theneuromuscular endplates. However, such restorativeevents are limited to less complex and unilaterallyoriented pathways, like the innervation of a singlemuscle or the arrangement of nerves with topologi-cally and functionally distinctive targets.

In spite of the ability of peripheral nerves to re-generate, surgical repair of nerve damage is not aseasy or even not possible for several reasons:

. the nerve is not accessible to surgically join theproximal with the distal stump,

. some neurons die inevitably due to hereditary

neuromuscular diseases,. the axons cannot be guided through the proper

pathway for their regeneration, and/or. regenerating axons do not nd their targets for

synaptic connections.

Diculties in surgery apply especially to the CNSwhere most surgical interventions are not possiblewithout destroying neighbouring parts of the nervetissue, where the network-like highest complexity ofmultiple circuits has been formed during develop-ment. In this network, molecular guidance does notproperly work any more, and oligodendrocytes seemto form an environment which inhibits axonalgrowth. Moreover, damaged neurons can be

destroyed by additional local mechanisms, probablydeveloped to eciently remove the sick elementsand preserve the remaining structural and functional

integrity of the tissue. Finally, astrocytes are respon-sible to ll the structural gaps with proliferation andto communicate will all other elements, thus balan-cing the decits.

Frequently, whole organs or part of the body areremoved or destroyed, e.g. a limb is lost which can-

not be replaced naturally, ultimately demanding anexternal intervention. In such cases, rudimentaryorgan of body functions (stability, standing up,walking) can be in part compensated with externalmechanical prosthetics. However, they are not act-ing on the neural function, but replace the periph-eral muscular, bone or joint functions. It seemsdesirable, but only in extremely rare cases possibleto date, to develop prosthetic interfaces between thenervous system and its various peripheral targets.These diculties are multiple and require rst theprofound understanding of the neuronal circuitryand function and second a directed ow of neuronalinformation from its natural origin through a sen-sing prostheses into a nal target in a functionally

remodelled form.A classical injury resulting in devastating func-

tional impairments is the incision of spinal cordwith usually combined lesions of ascending and des-cending paths, in addition to local necrosis ofintraspinal neurons. One of the predominant eldsof prosthetic intervention in the future will be thecure of paraplegic and paralytic individuals withbest results until now towards to reconstruct thesimpler reexes like that of the urinary bladder func-tion. Other work is carried out in order to stimulatemuscles of the legs and the back in order to makepossible getting up, standing or even walking. Firstattempts to achieve regeneration of neural connec-tions through regenerating nerves in the spinal cord

have been made, but they are still far away fromrestorative applicability.

Smaller, well integrated sensory organs and com-partments may be more accessible to so-called arti-cial sensors where a direct connection betweenelectronic devices and nerves is needed. The prob-ably most popular application where dierentdevices work benecially is the cochlear implant.Conceptually, further sensory organs may bereplaced by miniaturised and adaptable devices too.However, their success is limited up to date,although various concepts have been developed, e.g.to replace the retinal function with subretinal orepiretinal implants. In addition, further approacheswere devoted to replacing the optic nerve with per-

ipheral nerve implants, with limited success inregaining function, but encouraging results towardsunderstanding the mechanisms of axonal regener-

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ation and opening a new eld of natural neuopros-thetics in combination with articial prostheticdevices.

It is desirable, however, that all articialapproaches are accompanied by endeavours to fa-cilitate the natural approach, i.e. the regeneration ofaxons towards their natural targets and fresh for-mation of new synapses leading to at least partialfunctional recovery.

The goal of this work is to critically review someaspects of these approaches, particularly therequired functional abilities of implants, the designof microelectrode arrays (MEAs) for optimal func-tion and bio-(neuro-)compatibility, their biologicallyproper implantation and their function for recordingof nerve signals and/or stimulation of nerves and/ormuscles.

2. IMPLANTABLE ELECTRODES

2.1. General Remarks

The nervous system is the most complex biologi-cal system and is therefore characterised by highvulnerability to external violence and genetic dis-turbances. Profound knowledge about its develop-ment, functional consolidation and structuralorganisation is prerequisite for any attempt to estab-lish good recording or stimulation in order to treatdefects in a proper way. A number of proposedideas for electronic implants is based on a more ``en-gineer-like'' way of thinking than considering the

complexity of the biological systems. For instance,nerve bres are not ``wires'' or ``cables'' in the mech-anical point of view, but long, sensitive structuresconsisting of biological membranes with multiplereceptors and delicate interactive elements for on-line sensing the environment and transmitting infor-mation via molecules, potentials and changes in thechemo-electrical activity. The principal requirementsto any implantable structures are therefore featuresmimicking some of the biological functions of nervesand replacing these functions depending on thescope of implantation.

Many eorts have been undertaken to interfaceelectronics to nerves, and there is a reasonable num-ber of successful applications with dierent goals.

The aim of using implants in basic research is tounderstand working principles of the brain, to ana-lyse processing of information, to unravel principlesof connectivity within the nervous system and tostudy cellcell interactions in subsets of neuronalpopulations. Another goal is to replace functionswhich are lost due to damage of the nervous system.In this context, some attempts are concentrated onsensory functions and stimulation of muscles byimplanted electrodes and on development of closed-loop ``neuronal prostheses''. Microtechnology andmicroelectronics have obtained impressing capabili-ties which have to be combined with possibilities ofmodern biology and medicine under careful con-sideration of possibilities and constraints of neuroa-

natomy and neurophysiology. In fact, implantationof electrodes for nerve signal recording and nervestimulation has been carried out for some couples of

years with success, and there are also some promis-ing concepts for the future and interestingapproaches for special applications. A review about

the basics and dierent aspects of neural prosthesesis given by Agnew and McCreery (1990a).The success of neural prostheses depends basically

on their capability to record nerve signals and/or tostimulate nerves and muscles. It is obvious that suchimplants must full very special requirements whichare directed to the electrodes and the substratewhich carries the electrodes. Prerequisite of attach-ment to nerves and sustaining there is the biocom-patibility for nerve-specic implants, or more strictlyspoken, their neurocompatibility as shall beaddressed later.

The substrate which carries the electrodes must becompletely insulating to prevent cross-talkingbetween them. The choice of the material will also

depend on the possibilities and restrictions arisingfrom the fabrication process. In microelectronics,silicon is most frequently used. However, in the ma-

jority of implantation cases, it will be of favour touse exible material for the implants in order tomimic biological tissue and to reduce the possibilityof mechanical damage. As one of these exible ma-terials, polyimide obtained attention in the lastyears.

As they are directly in contact with the nerve ofinterest, the appropriate construction of theimplanted electrodes is of key importance for thesuccess of the whole articial system. They have tobe of a size comparable with the size of the neuronsin order to interact only with few or even one neur-

on. If electrodes are intended to interact with nervebres, i.e. axons, in the case of myelination their``radius of action'' has to be sucient to reach thenearest node of Ranvier. The small size requiresoptimised geometry and electrical properties for suf-cient recording ability and a high relative chargetransfer capacity. Moreover, all components have tobe adapted to this, i.e. electronic units are neededwith a high sensitivity, a high signal-to-noise ratio,and sucient shielding between single channelswithin the device and against external interferences.

2.2. Electrodes Implanted Without Nerve Cut

2.2.1. Cu Electrodes

The term ``cu electrodes'' applies to thosedevices which engulf the entire circumference of anerve. The shape of these electrodes or of the arrayof electrodes has to be adapted to the naturalarrangement and thickness of neurons or axons byalso taking in account the vascularisation at the siteof implantation.

Cu electrodes are applied preferably at periph-eral nerves. The advantage of cu electrodes is thatimplantation is relatively easy and that the nerve isnot damaged by proper surgical implantation.Possible displacement along the nerve bundle can becircumvented by mechanical xation at the site ofinterest. First models carried only one or two elec-

trodes. They were made using a platinum foil(``split-cylinder'', Avery, 1973) or platinum wiresembedded into insulating material (``wrap-around'',

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Hagfors, 1972). The rst models have been rathersti, often resulting in damage of nerves due tomechanical displacement, or pressing the nerve

bres, or disturbing the vascularisation and causingischemia. Modern cu electrodes try to avoid thisby the introduction of exible materials and adapt-able geometries, like a helix-shaped electrode(Agnew et al., 1988) or a spiral-cu electrode(Rozman, 1991) which allow adjustment of theimplant to the actual diameter of the nerve brebundle. Another possibility is the application of so-called ``half-cu'' electrodes rstly described by Kimet al. (1983) and patented by Testerman andBierbaum (1994). A particular design is the arrange-ment of exible interdigitating sub-units with micro-electrodes along a backbone-like carrier (Klepinski,1994; Meyer et al., 1995). Models of these two de-signs are shown in Fig. 1.

First models of cu electrodes did not completelyfull the requirements for accurate measurements,because they only allowed recording of supercialsum potentials with the axons in the centre of thenerve cylinder to contribute less signicantly to themeasured signal. In accordance, the inner axonswere less aected by stimulation than the supercialaxons of the bundle. For this reason, advancedmodels of cu electrodes have been developed withmore smaller electrodes arranged around the per-

imeter of the bre bundle which allowed toapproach central axons too. Topographic stimu-lation of nerve bres within a given nerve bundle

may sometimes be important, because dierentmuscles may be innervated by the bres of selectivelocalisation within the bundle, and often a mixtureof eerent and aerent bres occurs in the nerve.Such mixed populations may lead to undesirablesensations during stimulation. In the case of severalnerve bres innervating dierent muscle bres ofone muscle, cyclic stimulation can be performedwhich guarantees overall stimulation frequency,while stimulation frequency and thus fatigue of asingle muscle bre can be low (Happak et al., 1989;Talonen et al., 1990). With a sucient number ofelectrodes within the cu, high selectivity of stimu-lation can be achieved by either longitudinal ortransversal currents produced by appropriately

switched electrodes (Veraart et al., 1993). However,there is still the problem of possibly dierent brediameters which leads to dierent activationthresholds of the bres. Goodall et al. (1996) foundthat large bres were activated before smaller with acu electrode containing 12 electrodes arranged infour longitudinal tripoles, irrespective of bre pos-ition. Position selectivity could be enhanced by ahigher ratio of transversal to longitudinal currents.Nevertheless, the main prerequisite for successful

Fig. 1. Models of two recent possibilities of the design of cu electrodes. Left row: A ``half-cu'' elec-trode which surroundes the nerve and is secured with suture. Dots represent microelectrodes which are

placed along the cu and allow selective recording and stimulation (Kim et al., 1983). Right row: A ex-ible interdigitating cu electrode (``FLIC'') where microelectrodes are placed along ``ngers'' which bend

to form a tube for the nerve (Klepinski, 1994; Meyer et al., 1995).



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application of such sophisticated electrode designsand stimulation protocols is detailed knowledge ofthe distribution of nerve bres within the nerve bun-

dle and their function. Moreover, a really reliablexation of the cu must be achieved because other-wise the cu may rotate around the nerve or shiftalong the nerve, both leading to loss of selectivity.

2.2.2. Penetrating Electrodes

The so-called penetrating electrodes are anothertype of electrodes aiming to directly approach nervebres situated deeper in the tissue. The rst pene-trating electrodes were simply thin metal wires orneedles which were inserted into the nervous tissue.They often were insulated except a region on the tip.The other possibility was to use glass micropipettespulled out to very small inner diameters and lled

with saline. They also can be inserted into nervoustissue, and signals are recorded by a thin metal wireelectrode within the pipette. For simplicity, butwrongly, the glass micropipettes themselves oftenare designated as ``electrodes''. One recent examplefor the use of such pipettes is given in the work byWelsh et al. (1995), who investigated the role of theinferior olive, a major cerebellar aerent, by record-ing the activity of the Purkinje cells. For this pur-pose, they positioned up to 39 glass micropipetteswith electrodes (lled with saline, 12 MO, 24 mmtip diameter) independently 100125 mm below3 mm2 of the cortical surface of rats, with a distancebetween the electrodes of 250 mm. This procedure isreally time-consuming, and its application is limited

by its complexity. The animals have to be anaesthe-tised and held in a xed position. Moreover, theposition of the electrodes is not optimised withrespect to the location of the desired target, in thiswork the Purkinje cell layer, but recorded signalsare used from those electrodes which are placed atbest by chance.

Wedge-shaped microprobes carry a line of elec-trode sites for recording and stimulation and maytherefore be designated as ``one-dimensional'' arrays[Fig. 2(A), BeMent et al., 1986; Ensell et al., 1996;Kewley et al., 1997]. Typical dimensions are in therange of 2 mm for the shank length, 100 mm for thewidth and 20 mm for the thickness, but there arealready smaller structures in development. An on-

chip electronic circuitry would allow pre-amplica-tion and rst processing, but it is not yet realised inmost cases. Howard et al. (1996) built a recordingarray where they combined high-impedance micro-electrodes with low-impedance EEG electrodes, andactivity of human cortical neurons could berecorded.

For signal recording and stimulation over a largerarea, e.g. areas of the cortex, two-dimensionalMEAs are necessary as they are known as solid pla-nar arrays for in vitro experiments (Wilson et al.,1994; Nisch et al., 1994; Gross et al., 1995; Bove etal., 1997). A exible planar MEA with polyimide assubstrate and 24 gold microelectrodes (4040 mm,210 mm spacing) was made for recordings of electri-

cal activity of the cortex (Owens et al., 1995).Improvement of recording structures was achieved

by silicon microtechnology which made it possible

to create not only planar two-dimensional MEAsbut also penetrating electrode structures for in vivomeasurements. The substrate carrying the electrodesis either needle- or wedge-shaped to allow pen-

etration of the nervous tissue which makes possiblerecording from and stimulation of axons not onlyon the surface but also in a well-dened depthwithin the tissue, e.g. within the bre bundle orregions of the brain. Implanting such a device isalso associated with an accidental damage of the tis-sue, i.e. some neurons will be destroyed, and a cer-tain portion of the axons will be disrupted.Moreover, stiness of many models may lead todamage of nervous tissue. That is why the eortsare directed to miniaturise the penetrating parts ofthe implant and to use more exible materials.

Nordhausen et al. (1996) reported about a silicon-based two-dimensional MEA shaped as a grid of1010 needle electrodes with a spacing of 400 mm.

The needles are approx. 80 mm wide at the base and1.5 mm in length. They are insulated with polyimideexcept for approx. 50 mm at the tip, the latter beingcoated with platinum to form the active electrode.The electrode array was successfully applied for therecording of local visually evoked responses in thevisual cortex of cats at sub-sets of 15 electrodes(Nordhausen et al., 1996), and these measurementswere continued in order to rene recording and sig-nal processing procedures (Maynard et al., 1997).

A similar approach was performed by Rutten etal. (1995). They created a three-dimensional needlearray with 128 recording sites with one electrode onthe tip of a needle. The needles are made from sili-con and are embedded into a glass substrate. They

vary in height from 250 to 600 mm and have a dis-tance of 120 mm, with a tip size of 1515 mm. Thedierent length of the needles should allow reaching

Fig. 2. Dierent layouts of penetrating electrodes. (A)Wedge- or shank-shaped electrode carriers are one of themost common designs developed and applied by severalgroups. In many cases, more than one shank are combined.

Typically, the shanks are 13 mm long, 30100 mm wideand 820 mm thick. (B) Proposed structure of an array ofneedles carrying multiple electrode sites thus reaching areally three-dimensional arrangement. (C) A combinationof cu and penetrating electrodes (Durand and Tyler,1995; Tyler and Durand, 1997). Only one possible designof the implant is shown in opened and closed states. Theelectrodes penetrate the nerve when the cu is bent aroundthe nerve. (D) Flexible nerve plate (Meyer et al., 1995)which can be inserted into nerve tissue, e.g. the retina or

the cortex, or into a peripheral nerve fascicle.



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approximately three-dimensional areas of nervoustissue, e.g. in a peripheral nerve fascicle.

It could be suggested that the optimal MEA

would be one with the possibility to move eachsingle electrode separately to its best place in orderto optimise recording or stimulation. On the onehand, this is limited by the possibilities of microtech-nology, because it would be very complicated to de-sign and fabricate a MEA with independentlymovable and mechanically stable electrodes withperfect conduction of signals and insulation in theaqueous surrounding. On the other hand, there isthe biological issue to nd out which place withinthe nervous tissue is actually the best for the desiredpurpose. This is time consuming, and it is dicult topredict how much of the nervous tissue is damagedduring penetration of the electrodes and their lateralmovement.

In view of these problems, it would be useful tocombine the approaches mentioned above and cre-ate a three-dimensional MEA by arranging needleprobes in a grid with electrodes placed along theprobe shank as shown in Fig. 2(B). This would pro-vide a high-density three-dimensional arrangementof electrode sites, and best of them could then befound out after successful implantation by testingrecording or stimulation characteristics.

Of course, there is a huge number of technicalproblems associated with the very small dimensionsof MEAs and also of the design shown in Fig. 2(B).One of the problems is that, on the insulating sub-strate, conducting electrode sites and leads must beplaced. A second problem is that cross-talk between

the leads and leakage have to be avoided. A thirddrawback is that with increasing number of electro-des the number of leads increases too, whichdemands an intricate on-chip design and reliableconnections and cables to electronic processingunits. Furthermore, the whole set-up must operatereliably over a long time period. Last but not least,processing of a high number of channels requires asophisticated microelectronics which may not be toobig and energy-consuming for practical every-dayuse.

With the technological advance in the eld of con-structing exible polymeric substrates, so-called`` exible nerve plates'' come into development[Fig. 2(D), Meyer et al., 1995]. They consist of a

exible substrate which carries a MEA and possiblyother elements of a microelectronic circuitry. Theexible nerve plate could be applied in a big varietyof tissues and organs. For instance, it could beinserted into the retina or longitudinally between thebres of a nerve bundle or a muscle, and act there-fore as a kind of penetrating electrode in thesecases, or it could be placed onto the surface ofspecial regions of the cortex depending of the scopeof the experiment. The exibility of the nerve platewould allow better protection of nerve tissue andvery short distances between electrodes and neuronsor bres which are necessary for an optimum decod-ing of nerve signals or transmitting stimuli to thenerves.

As for cu electrodes, the geometry and thearrangement of penetrating electrode array vari-ations are currently optimised to t with geometrical

nerve bre characteristics (bre position in a bundle,diameter of the bres and their orientation).Inuence of specic anodecathode combinations

and stimulus parameters are also under research.There are some research projects dealing with thedevelopment of a computer-assisted control of sti-mulating potential pulses for the formation ofspatially distinguished electric eld within the ner-vous tissue which allow selective stimulation of thedesired neurons. By sophisticated control of theheight and time of applied potential pulses, a morediscrete stimulation of the axons can be achieved.Such eorts are undertaken for electrode arrays forboth central and peripheral nerve systems.

2.3. Electrodes for Regenerating Nerves

A dierent approach is performed by the so-called

``electrodes for regenerating nerves''. They aredesigned as a MEA placed on a sieve-shaped (i.e.perforated) plate which contains holes which can beround or rectangular or even shaped as long narrowslots. The microelectrodes are situated nearby theholes or are part of the hole's wall in order to opti-mally record or to stimulate. The principal idea canbe described very briey: The nerve is cut rstly,then the electrode array is adapted into the expectedpath of the regenerating bres in a fashion that thenerve bres are allowed to regenerate through theperforations of the device (Fig. 3). The distal stumpof the cut nerve is aligned at the opposite side of theelectrode array in order to be used by the axonsexiting from the perforations as a guidance path for

further growth. In most cases, the perineural sheathof the nerve can be replaced in the area of insertionwith polymeric tubes which act as mechanical stabil-isers.

The advantage of this approach is that with thisdevice the electrodes are in intimate contact with thenerve bres, this allowing both accurate recordingand ecient stimulation. Both procedures areexpected to be performed relatively reliable becausewith a proper choice of the hole diameter a predict-able number of axons would regenerate through in-dividual perforations. Moreover, the micro-electrodes remain always in the same position rela-

Fig. 3. Principal concept of regenerative electrodes. Axonsregenerating from the proximal stump of a dissected nervegrow through a permissible array of electrodes carried by asubstrate. The array may be designed as a kind of mesh or

sieve. Latticed arrangements are also possible. The aper-tures (holes) of the substrate determine and x the position

of the regenerating axons relative to the electrodes.



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tively to the nerve bres because the bres are xedby the holes they grow through.

The obvious drawback of this method is that the

nerve has to be cut in order to regenerate throughthe implanted device. The success of the whole oper-ation can be evaluated only several weeks later,when the axons have regenerated through thedevice. The second disadvantage is its applicabilityonly in peripheral nerves, because central nervepathways do not regenerate spontaneously. Thethird disadvantage is that the device limits the regen-eration of some neurites, namely of those, whosegrowth cones hit on the device and fail to elongatewithin one of the pre-drilled holes. In spite of theselimitations, such electrodes display an elegant wayof application in the peripheral nerves.

Indeed, attempts to develop and utilise such elec-trodes have been performed continuously since the

sixties (for an overview, see Kovacs and Rosen,1992). Mannard et al. (1974) combined 10 silverwires in a conical bundle. They were carried by aattened epoxy bulb where holes have been drilledin. Llina s et al. (1973) proposed a relativelyadvanced concept for an electrode device with aradial array of gold electrodes surrounding theholes. Loeb et al. (1977) reported on the fabricationof an electrode array with 0.31.2 mm long tubesthe regenerating axons should be growing through.Edell (1980) developed a structure with long narrowslots for nerve bre regeneration. Electrodes wereplaced on the thin interspaces between the slots.Rosen and Grosser (1986) also proposed a conceptof ``regenerative electrodes'' in order to ``restore

normal nerve impulse communication and hencenerve function''. A micromachined silicon electrodewas made by Akin and Naja (1991). After implan-tation, nerve regeneration through this device couldbe achieved with the glossopharyngeal nerve of ratsas model system (Bradley et al., 1992), and nervesignals could be recorded (Akin et al., 1994), evenover a longer time period (Bradley et al., 1997). Acomplete set-up of a perforated electrode with aMEA was also developed and applied by Kovacs(1991). After implantation into the peroneal nerve inthe hind limb of rats, some spontaneous action po-tentials could be recorded (Kovacs et al., 1992).Later on, more extended recordings from the ratperoneal nerve and the bullfrog cranial nerve were

reported (Kovacs et al., 1994). The two last devicesare shown in Fig. 4. Recently, a perforated MEAwas described which was implanted between the cutends of the rat sciatic nerve (Navarro et al., 1996).Another description of this MEA was given byDario et al. (1997).

The biological and technical aspects concerningthe electrodes with perforating paths are complexand do not permit a simplication in use, neither ageneralisation in terms of their applicability, mainlydue to the traumatic procedure of implantation.

Although techniques of microfabrication havebeen developed rapidly during the last years, someproblems arise if a large number of microelectrodeswith the accompanying connections has to be placed

and xed on a silicon chip. These problems areenhanced in electrodes for regeneration, because theholes take a big part of the total surface area and

limit per se the number of connections with individ-ual holes. The shortage of surface of a microchiplimits both the number and the size of holes for agiven electrode.

The minimum requirement for a perforating elec-trode is that growth cones t within the individualhole and pass through its lumen without disturb-ances in their growth properties, and later on, intheir conduction velocities. It does not seem criticalwhether the holes are round or rectangular.Experiments of dierent groups showed that,depending on the nerve where the chip is placed, theminimum size of a hole should be around 25 mm,

and thus in the range of the thickness of a singleaxon. Some groups use hole sizes between 25 and50 mm and expect that more than one axon cangrow through each hole. The natural intention of aquantitatively high performance of regenerationrequires as many small or larger holes as possible.Higher density of holes limits, on the other hand,the free space for connecting the electrodes on themicrochip. A compromising way to use fewer con-nections with more bioelectronic interfaces is themultiplexing of the electrodes, although this processcannot be forced endlessly. Thus, the scope of theexperimental design in individual cases determinesthe ratio between number of perforations and num-ber of electrodes, again depending on the cross-sec-

tional area of the nerve to be used. As an example,a ratio of 2.5:1 between the area of the chip bearingthe holes and the cross-sectional area of the nerve is

Fig. 4. Electrodes for regenerating nerves as made by (A)Akin and Naja (1991) and (B) Kovacs et al. (1991).Redrawn from original papers with permission from the

authors and IEEE.



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3. BIOCOMPATIBILITY

3.1. General Remarks

The major prerequisite for the application ofimplants, e.g. neuroprostheses, is that the organismaccepts the implant, i.e. that the implant is biocom-

patible. It is widely accepted to dene biocompatibil-ity as ``the ability of a material to perform with anappropriate host response in a specic application''(Williams, 1987). This broad denition comprisesaspects of biological, chemical and physical proper-ties of the implant which will be addressed in thischapter. Nevertheless, the concept of biocompatibil-ity is still disputed and depends strongly on the ap-plication eld of implants. Whereas in some casessurface (chemical) composition of the implant is animportant parameter, in other cases physical proper-ties (size, shape, stiness) are the major determi-

nants of biocompatibility, particularly under theinuence of locomotion (Boss et al., 1995).

An implantation is always a traumatic interven-tion. However, one important way to minimise theconsequences is a high biocompatibility of theimplant. An implant can be considered to be bio-compatible if

. it does not evoke a toxic, allergic or immunologicreaction,

. it does not harm or destroy enzymes, cells or tis-sues,

. it does not cause thrombosis or tumours,

. it remains for a long term within the organismwithout encapsulation or rejection.

For a long-term stable neuroprosthesis, the wholeimplant must have mechanical and geometricalproperties which t to the site of implantation inorder to minimise traumatic lesions. The concept ofbiocompatibility is not limited to non-toxicity, butencloses physical and chemical surface propertiesand whole behaviour of the implant in its biologicalenvironment as well. Therefore, two areas have tobe considered, the ``biosafety'' and the ``biofunction-ality''. Biosafety means that the implant dos notharm its host in any way, and biofunctionalitymeans that the implant acts in the body as it wasintended. In addition, ``biostability'' is importantwhich means that the implant must not be suscep-

tible to attack of biological uids, proteases, macro-phages or any substances of the metabolism. Forexample, implants may be subject to continuousattack by hydrolytic enzymes (Salthouse, 1976) orfree radicals produced by monocytes and/or celllysis. Stability of implanted material is importantnot only for stable function, but also because degra-dation products may be harmful for the host organ-ism. Overviews about biological reactions toimplanted materials have been given, e.g. by Henchand Ethridge (1982), Anderson (1988), Tang andEaton (1995), and Ratner et al. (1996).

Once implanted, a neuroprosthesis has to remainwithin the body of the patient for many years. Insome cases, it is intended to cover the whole life of

the patient. This is in particular obligatory inimplants for regenerating nerves, because theywould have to be cut again in order to remove the

implant. This implies extreme demands on stabilityof function and set-up of the implant. Where necess-ary, e.g. in muscle tissue or in joints, wires and sub-

strates have to be exible enough to allow multiplebending without damage of surrounding tissue andwithout breaking at the end. Sharp edges whichcould damage cells and tissues or rough surfaceswhich allow attachment and growth of microbeshave to be avoided. No corrosion is allowed, andinsulating polymers must keep their insulating prop-erties. The implant must remain in its implantationsite, thus a reliable xation must be guaranteed, e.g.by appropriate suturing. In addition, used polymersmay not release any substances, e.g. monomers oroligomers, modiers, sterilising agents like ethyleneoxide, etc. In addition to the implant itself, also ma-terials used for the xation of an implant must bebiocompatible, as addressed in the perspectives of

brain implants by Mod et al. (1997).The exception where release of substances isintended are so-called drug delivery systems.Antibiotics, hormones and growth factors could bereleased in order to prevent sepsis and to improvewound healing, tissue repair and nerve regeneration.The materials of these systems also must be biocom-patible, with particular emphasis on preventing pro-tein adsorption and platelet adhesion which couldhinder the substances to be released (Park and Park,1996).

3.2. Implant Materials

It is obvious that requirements towards implantsare very high. Nevertheless, there are many ma-terials (metals and polymers) which meet theserequirements at least to a big portion.

Nowadays, platinum is the electrode material ofchoice, because it is stable and inert. The amount ofplatinum ions released into the surrounding tissuemay be neglected even after long terms of stimu-lation. During the last years, iridium has been ofincreasing importance because a stable oxide lmcan be formed on the surface of iridium electrodes.This oxide lm has a big charge delivery capacity

and is, for this reason, well suited for stimulatingelectrodes. Carbon bres or glassy carbon are alsoused as electrode materials, and they are biocompa-tible and stable, though they have a higher rough-ness than metals. Platinum and iridium areestablished materials in microelectronics, and alsocarbon can be deposited onto microelectronic struc-tures.

As recently reviewed by R hova (1996), polymersare used as carrier material and for encapsulationpurposes. Most common materials are epoxy resins,polytetra uoroethylene (PTFE, Teon1), siliconerubbers and polyimide. These polymers are biocom-patible, electrically insulating and stable. The bulkproperties of the polymers can be modied to a cer-

tain degree, and also surface modication pro-cedures are performed in order to improvebiocompatibility.



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3.3. Response to Implantation

When an implant is brought into the body, therst event is that proteins adsorb onto the surface of

the implant. A dozen proteins can be found in bio-logical uids at concentrations higher than 1 mg/ml,and certainly they will form major parts of layersformed on the implant at least in the initial state ofadsorption (Andrade and Hlady, 1987). The detailsof this process depend on the surface of the implant,the composition of the biological environment andthe nature of the adsorbed proteins. Adsorption ofproteins such as collagen or bronectin can favouradhesion of tissue cells (Seeger and Klingman, 1988;Drumheller et al., 1994). An encapsulation of theimplant by autologous material (astrocytes, proteinlayers, endothelial cells, broblasts) is desirable inorder to integrate the implant into the organism and``mask'' it in order to avoid undesired reactions ofthe immune system, thus promoting incorporationand acceptance of the implant. On the other hand, itwas reported that bronectin and brinogen canenhance the adhesion of dierent bacteria, e.g.Staphylococcus aureus, which is a common cause ofinfections after implantation (Vaudaux et al., 1984,1995). Adhesion of the uropathogen Pseudomonasaeruginosa B4 onto polystyrene was increased whenurine-derived a-1-microglobulin was pre-adsorbed(Wassal et al., 1995).

Inammatory reactions can be another conse-quence of an implantation. Inammations involvevascular, neurological, humoral and cellular re-sponses, and the following acute-phase response is

characterised by stress-induced changes in theneuroendocrine and immune systems (Kushner,1982; Khansari et al., 1990). During an acute-phaseresponse, concentration of dierent serum plasmaproteins increases signicantly, such as serum amy-loid A, C-reactive protein, brinogen, a-1-antichy-motrypsin, complement protein factor B and C3,haptoglobin and a-1-antitrypsin, whereas concen-tration of other proteins is decreased (albumin,transferrin and transthyretin) (Baumann andGauldie, 1990). Furthermore, leukocytes adhere toblood vessel walls and migrate into the tissue whichrequires highly specic interactions mediated byselectins, integrins and inmmunoglobulins (Jones etal., 1996). Cytokines are released, e.g. IFN-g, TNF-

a, IL-1 and IL-6, which are principal mediators ofthe acute-phase response (Baumann and Gauldie,1990). High concentrations of TNF-a and IL-6, e.g.may induce tissue damage, up-regulation of surfaceadhesion molecules and enhanced production ofproteases and free radicals by macrophages.Another issues of an in ammation are fever andinltration of monocytes/macrophages, eosinophils,neutrophils, lymphocytes, granulocytes, broblastsand giant cells.

When an implant is brought into the brain, astro-cytes can be observed to respond quickly to thisinjury as they react to every damage (Cavanagh,1970; Ludwin, 1985). They proliferate in the vicinityof the implant and send their processes towards the

implant. Microglial cells are also activated andtransform from ramied to amoeboid type. Theimplant is coated with a layer which contains col-

lagen, giant cells, nerve bres in dierent states andblood vessels (Schultz and Willey, 1976).

In order to evaluate biocompatibility of a ma-

terial, in vitro experiments play the major role(Klein et al., 1995; Hanks et al., 1996), althoughthey do not reect the whole complexity of the invivo situation. In vitro experiments are based on celllines and on primary cell cultures depending on theintended site of application. After bringing the cellsin contact with the material, dierent parametersare evaluated, as morphological and ultrastructuralchanges, cell adhesion, release of mediators, presen-tation of molecules, changes of metabolism, etc.One of these parameters is metabolism of arachido-nic acid by macrophages as indicator for inamma-tory processes (Charissoux et al., 1996). Release ofcytokines is also a measured parameter, e.g. ofTNF-a (Hunt et al., 1996). Although neuroprosth-

eses are not applied inside blood vessels, blood com-patibility is also an important issue, with particularemphasis on behaviour of blood cells, extent ofcomplement system activation, platelet activationand clot formation. The ability of endothelial cellsto produce antithrombotic substances has to beexamined (Cenni et al., 1993).

There are dierent cell surface molecules whichare important in the context of implants.Endothelial cell cultures are preferred for biocom-patibility studies of materials for vascular pros-theses, and proteins important for adhesion areexpressed upon contact between the implant and thecells. The occurrence of such proteins is studied in

biocompatibility evaluations with endothelial cells.Examples are the platelet endothelial cell adhesionmolecule-1 (PECAM-1, CD31), the endothelial leu-cocyte adhesion molecule-1 (ELAM-1), the intercel-lular adhesion molecule-1 (ICAM-1, CD54), and thevascular cell adhesion molecule-1 (VCAM-1) (Cenniet al., 1995).

One aspect of biocompatibility for neural pros-theses is damage by permanent charge injection.Permanent electrical stimulation can cause damageof neural tissue, such as gliosis, calcication of neur-ons and other cells, lipid inclusions, or glycogengranules in astrocytes, and neural tissue may be lostnally (see Agnew and McCreery, 1990a). The basicparameter for stimulation intensity is the relationbetween charge density and injected charge perphase. Charge density is measured in mC/cm2 and isdetermined by the frequency, the applied currentand the size of the electrode. Absence of neuraldamage can be expected for the range of a chargedensity of 100 mC/cm2 with 0.2 mC/phase or a chargedensity of 15 mC/cm2 with 8 mC/phase (Agnew andMcCreery, 1990a, p. 229).

Agnew and McCreery (1990b) came to the con-clusion that damage of neurons and axons has itsorigin in their hyperactivity due to severe electricalstimulation, particularly if stimulation is performedwith a high frequency. That is why frequency ofstimulation should be as low and pulse duration

should be as short as possible. Furthermore, stimu-lation should not be performed continuously butwith cut-os where possible.



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3.4. Surface Modication

In order to enhance biocompatibility of implantedmaterials, reduce macrophage adhesion onto the

implants and prevent inammatory reactions, sur-face modications of materials intended for implan-tation are widely studied. These investigations areperformed particularly with polymers because theyare the main material used for housing, encapsula-tion and insulation purposes. Whereas attractionand adhesion of macrophages and other whiteblood cells to an implant can favour inammations,inhibition of such an adhesion would be an import-ant factor of biocompatibility.

Nevertheless, it depends on the intended use ofthe implant whether adhesion of proteins and cells isdesirable. The best example are vascular graftswhich require a completely dierent cell behaviouron their surfaces: inside the grafts minimal cellular

adhesion and brin formation, but extensive ad-hesion of tissue cells and matrix formation on theouter side. With a neuroprosthesis, also dierentproperties are needed. Recording electrodes shouldremain bare for good sensitivity, and the best wouldbe an intimate and stable contact between the elec-trode and neuron cell bodies or axons. Housing andencapsulation materials would be allowed to becoated by tissue cells like broblasts in order to in-corporate them into the body. Stimulating electrodesare not coated if the applied currents are highenough to cause Faradaic processes at their surfacelike oxygen evolution, and processes of in am-mation, gliosis and neuron damage occur not actu-ally on the electrode surface, but nearby theelectrodes.

3.4.1. Surface Roughening

First kind of surface modication which can bedone is the preparation of smooth or rough surfaces.If minimal cell adhesion is intended, surfaces aremade as smooth as possible. If cell adhesion iswished in order to achieve good incorporation intothe body, rough surfaces are advantageous.Roughening is performed by particular productiontechniques, where roughness is an intrinsic propertyof the material, or it can be done by micromachiningtechniques, where particular patterns are applied tothe smooth material, e.g. by photolithographic or

micromechanical procedures. In many cases, V-shaped grooves are milled or etched into the ma-terial in order to align growth of broblasts and tox the implant within the tissue. One importantexample are dental prostheses, where reliable attach-ment is essential for long-term clinical success. Longgrooves running perpendicularly to direction ofinsertion of dental implants could impede down-growth of epithelial cells on the implant, whichallows connective tissue to adhere and leads to a bet-ter xation of the implant (Chehroudi et al., 1991).

For many years, carbon has been used for dier-ent kinds of implants, such as dental implants, per-cutaneous devices, tendon and tracheal substitutionsand heart valves (Haubold et al., 1979; Iumashev et

al., 1983; Alberktsson et al., 1986; Tian et al., 1993).Carbon is known for its excellent biocompatibility,and it is also used for coating of polymeric materials

to improve properties of the latter (Pizzoferrato etal., 1993; Cenni et al., 1995). Main modications ofcarbon used for implantation purposes are pyrolitic

carbon and glassy carbon. Both modications havea turbostratic structure which leads to a microscopi-cally heterogeneous surface distribution of electrondensity and a microscopic roughness (Jenkins et al.,1972; Wege, 1984; Oberlin, 1989). One possiblereason for the observed good cell adhesion proper-ties is the combination of hydrophobic surfacewhich favours protein adsorption and roughnesswhich gives cells good ``points of attack'' for ad-hesion.

3.4.2. Chemical Modication

Another kind of surface modication is to changethe chemical composition. Ratner (1997) dividesthese surface modication procedures into biological

and non-biological methods depending on the kindof molecules bound to the surface:

. non-biological modication: functional groups(amine, hydroxy, carboxy), sulphonates, n-alkylchains, hydrogels, polymers such as poly(ethyleneoxide), poly(2-hydroxyethyl methacrylate), poly(n-vinyl pyrrolidone), polyacrylamide, and

. biological modication: coating with heparin,hyaluronic acid, sugars, peptides, lipids, enzymesor growth factors.

Naturally, layers created by dierent methods mustbe stable unless not being intended for biodegrada-tion. After their fabrication, no reactivity mayretain, for example by non-saturated binding sites,

e.g. sites activated with carbodiimide) or reactivedouble bonds (e.g. after cross-linking with alde-hydes).

The well-known statement that ``water is the mostbiocompatible substance we know'' would lead tothe conclusion that hydrophilic and therefore wetta-ble surfaces should be advantageous. Several sub-stances are hydrophilic per se, e.g. glass or silicondioxide. In other cases, hydrophilic surfaces can beobtained by polar or ionic groups. Such functionalgroups can be placed on the surface by chemicalmodication of surfaces or by coating with appro-priate molecules, e.g. with proteins or sugars. Witha polymer, monomers can be applied for the prep-aration which carry ionic or polar groups, or co-

polymers can be prepared in order to combinedierent properties for desired mechanical proper-ties and surface chemistry (see, for example, Kishidaet al., 1991; Arshady, 1993). A surface which has tobecome hydrophilic can have anionic properties byintroduction of negatively charged groups, particu-larly carboxy (0COO) groups. Neutral hydrophilicsurfaces can be obtained by hydroxy (0OH) oramide (0CONH) groups, and cationic hydrophilicsurfaces are made by introduction of dierentamino groups (0NH2, 0NHR0, or NR3).

Yun et al. (1995) modied PTFE surfaces withdierent functional groups and found that cell ad-hesion and cytokine release were inhibited at bestwhen modication was performed with amide

groups to obtain neutral hydrophilic surfaces.Coating of hydrophilic surfaces with polysacchar-ides, e.g. dextrans, has been recognised to be advan-



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tageous because each monomer within the chain car-ries up to three hydroxy groups. Indeed, non-specicprotein adsorption can be reduced (O sterberg et al.,

1993; Marchant et al., 1994).In order to prepare hydrophilic surfaces and toprevent undesired protein adsorption, surfaces werecoated with poly(ethylene oxide), heparin and albu-min, and a reduced thrombogenicity accompaniedwith a reduced plasma protein adsorption and plate-let adhesion was found (Amiji and Park, 1993). Theauthors explain these eects by a steric repulsionmechanism because the surface molecules can beregarded as entropic ``springs'' (Milner, 1991). Aspecial technique for the deposition of protein mol-ecules is cross-linking using glutaraldehyde or carbo-diimide as shown for arterial prostheses coated withcross-linked gelatine (Drury et al., 1987; Bordenaveet al., 1989; Marois et al., 1995), albumin (Guidoin

et al., 1984; Cziperle et al., 1992; Marois et al.,1996) or collagen (Noishiki and Chvapil, 1987;Guidoin et al., 1989). Another possibility for cross-linking is derivatisation of molecules to be depositedwith photodimerisable groups, such as thymine, cin-namate or coumarin. Upon irradiation with UVlight, these groups bind, and derivatised moleculesare deposited onto the surface, as demonstrated forthe deposition of chondroitin sulphate, hyaluronicacid and gelatine onto Dacron or polyurethane(Kito and Matsuda, 1996).

A dierent kind of introducing hydrophilicgroups into a surface is treatment with plasma in aglow-discharge apparatus (Yasuda, 1985; Moroso,1990; Piskin, 1992). The plasma can be made of

dierent substances, e.g. from oxygen, ammonia orwater, but also from organic molecules such asmethacrylates or siloxanes, and a big variety of ma-terials can be treated, even materials which are nor-mally inert, such as polystyrene or PTFE. On suchinert materials, plasma treatment is also used inorder to form reactive groups for further binding ofother molecules. Bigger molecules like polyethyleneglycols can also be deposited, and resulting layersresist protein adsorption and cellular attachment, ascould be shown for tetraethylene glycol dimethylether (Lo pez et al., 1991). Plasma polymerised silox-anes were shown to provide an anti-thrombogenicsurface which is important for vascular grafts andoxygenator devices (Hu et al., 1997).

3.4.3. Specic Modications

It has been said already that adsorption of pro-teins onto implant materials is desirable in thosecases where tissue cells may adhere in order to facili-tate wound healing and incorporation of theimplant. Adhering proteins serve as a kind of ``tem-plate'' to recruit healthy cells from intact tissuearound the site of damage. If the electrodes areimplanted into nerves, their appropriately designedsurfaces should incite the neurons to adhere to theelectrodes and to regenerate their axons along theelectrodes. First investigations for the purpose ofperipheral nerve regeneration were performed with

biodegradable polymers derived, e.g. from extractsof the extracellular matrix (Yannas et al., 1987;Aebischer, 1988; Chang et al., 1989).

For a rational design of regeneration-promotingsurfaces, it is necessary to nd out the key structureswhich give rise to the desired behaviour of cells, e.g.

neuron adhesion and axonal regeneration. Oneexample of such a key structure is the tripeptidesequence RGD (Arg-Gly-Asp) which is present inbronectin, a protein of the extracellular matrix(ECM), and to which many types of cells bind(Ruoslahti and Piersbacher, 1987; Hynes, 1990).Besides the big variety of binding cells, the import-ance of the RGD sequence is also underlined by thefact that it occurs in other adhesive proteins too,e.g. laminin (Grant et al., 1989), collagens, brino-gen, vitronectin, von Willebrand factor (Hynes,1987), entactin (Chakravarti et al., 1990), albolabrin,rhodostomin and other viper venom proteins listedin Soszka et al. (1991) and Chiang et al. (1996),tenascin (Sriramarao et al., 1993), and a zinc protein

(Takagaki et al., 1994). Thus, it has been possible toreduce the big protein molecule (molecular weightof 500,000) to a small tripeptide (molecular weightof 292) with the relevant function. Various materialswere modied by dierent methods with the RGDpeptide, as shown for glass, polyethylene terephthal-ate (PET) and PTFE by Massia and Hubbell (1990,1991), for polyacrylamide by Brandley and Schnar(1989), for poly(ethylene acrylate) by Hirano et al.(1993), for poly(g-methyl L-glutamate) by Kugo etal. (1994), for cross-linked polymer networks byDrumheller and Hubbell (1994), and for poly(vinylalcohol) by Sugawara and Matsuda (1995). Whereasthe non-modied materials showed only poor celladhesion, it could be observed in a high degree after

coating with RGD peptides, and also cell spreadingand migration were observed in some cases. This in-dicates that modication of implant surfaces with``biologically inspired'' synthetic molecules would beable to promote incorporation of the implants intothe tissue.

Other peptide sequences in proteins had been alsoidentied to be important for cell attachment. Inbronectin, six additional adhesion sequences havebeen found besides RGD which are listed inMooradian et al. (1993). In thrombospondin-1, thesequences RFYVVMWK and IRVVM were foundto support attachment of cells (Kosfeld and Frazier,1993). In the C-reactive protein, the adhesivesequence FTVCL was found (Mullenix et al., 1994).

However, there are only few cases of utilisation ofthese peptides. The bronectin-derived sequenceWQPPRARI was immobilised on polystyrene andPET, and enhanced adhesion and spreading of en-dothelial cells was found on the resulting surfaces(Huebsch et al., 1996).

Laminin is probably the most investigated ECMprotein. It is particularly interesting because it is anabundant component of the basement membranesduring development of the embryonic nervous sys-tem, but also present in the mature nervous systemwith apparently important functions not onlyrestricted to guidance or adhesion. In the developingand maturing central nervous system (CNS), lami-nin plays a crucial role, e.g. in cell migration, dier-

entiation and axonal growth (Martin and Timpl,1987; Kleinman et al., 1987; Martin et al., 1988).Besides the characterisation in vivo, it has been



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extensively used as a substrate for studies of thegrowth of neurons in vitro. Laminin is a big, multi-domain protein (Beck et al., 1990) with many bind-

ing sites for dierent cell receptors (Castronovo,1993; Mercurio, 1995; Gullberg and Ekblom, 1996;Wei et al., 1997). Amino acid sequences of lamininwhich have been identied to be important to celladhesion or growth are, e.g. CDPGYIGSR (Grafet al., 1987), RYVVLPRPVCFEKGMNYTVR(Charonis et al., 1988), SIKVAV (Tashiro et al.,1989), SRARKQAASIKVAVSADR (Sephel et al.,1989), RNIAEIIKDI (Liesi et al., 1989), PDSGR(Kleinman et al., 1989), CQAGTFALRGDNPQG(Tashiro et al., 1991), KQNCLSSRASFRGCVR-NLRLSR (Gehlsen et al., 1992), YFQRYLI(Tashiro et al., 1994), SIYITRF, IAFQRN andLQVQLSIR (Nomizu et al., 1995).

However, only few of these peptides have been

used for the modication of surfaces until now. Thepeptides used in most cases are YIGSR and IKVAVor longer versions of these sequences. Massia andHubbell immobilised GYIGSR onto glass (1990)and PET and PTFE (1991) and obtained adhesionand spreading of human broblasts. Hirano et al.(1993) coupled YIGSR and YIGSR-NH2 to poly(-ethylene acrylate) and obtained slightly enhancedattachment of dierent cell lines. In these cases,eects of YIGSR sequence were slightly lower thanthose of RGD. GYIGSRY was coupled to poly(-ethylene glycol) in a cross-linked network, and goodbroblast adhesion was achieved vs. no adhesion onbare polymer (Drumheller and Hubbell, 1994).Fluorinated poly(ethylene propylene) was modied

with YIGSR and an IKVAV peptide with 19 aminoacids (19-mer IKVAV), and attachment of neuro-blastoma and PC12 cells was evaluated (Ranieri etal., 1995). In most of these cases, it was reportedthat albumin pre-adsorption was necessary for e-cacy of immobilised peptides, probably becauseadsorbed albumin helped the bound peptide toachieve a relatively natural conformation on thehydrophobic surfaces.

Oenha usser et al. (1997) coupled the 19-merIKVAV peptide onto amine-modied glass, siliconwafers or microelectronic surfaces. Embryonic rathippocampal neurons attached to the modied sur-face and developed processes. It could be shown bythe patch-clamp technique that these neurons were

able to produce action potentials. Another lamininpeptide, RNIAEIIKDI, was used by Matsuzawa etal. (1996) to modify glass substrates. Again, embryo-nic rat hippocampal neurons attached to the modi-ed surface and developed a mature morphologywith outgrowing axons similar to that of neuronsgrowing on laminin. In both cases, a chemicallydened medium without serum was used. This maybe possible because the surface is hydrophilic due toits prior modication with amino groups.

In our laboratory, we investigate possibilities ofmodication of electrode surfaces in order toachieve a stable contact between neurons and elec-trodes. For this purpose, we tested dierent lamininpeptides (SRARKQAASIKVAVSADR and SIK-

VAV, RNIAEIIKDA, YFQRYLI, CDPGYIGSR,PDSGR, GTFALRGDNPQ) which were preparedby Kienle (1997). The immobilisation method

should be quick and simple, and, furthermore,should modify only the electrodes, but not the sub-strate without expensive photolithographic tech-

niques. These requirements directly lead to thetechnique of electrochemical polymerisation. Thepeptides were modied with a polymerisable group,3-hydroxyphenylacetic acid. Immobilisation of pep-tides by this technique could be demonstratedalready for antigenic peptides (Heiduschka et al.,1996, 1997). After immobilisation of the peptideslisted above onto glassy carbon, enhanced adhesionof embryonic chicken neurons was found, and neur-ite outgrowth from both adhered neurons andstripes of retina tissue could be observed (Huber etal., 1998). These eects diered depending on thepeptide, and laminin as whole protein moleculeshowed better ecacy when immobilised, particu-larly for the outgrowth of axons out of retina tissue.

Best results with synthetic peptides were obtainedwith the 18-mer IKVAV peptide. It further could beshown that electrochemically polymerised peptidelayers on recording electrodes which provide an inti-mate contact between neurons and electrodes wouldnot insulate the electrode. After polymerisation of 3-hydroxyphenylacetic acid, the impedance of micro-electrodes did not change signicantly (Valderramaet al., 1995).

For future developments, it would be desirable tond special peptide sequences to which each type ofcell responds specically upon contact with themodied surface. In this case, attachment and repul-sion, activation or deactivation of each type of cellcould be directed in the desired way as sketched in

Fig. 6.

4. AUDITORY IMPLANTS

The cochlea implant is one of the rst implantswhich have been developed for practical use, and ithas been working well since more than ten years inpatients suering from auditory impairment. It willbe therefore discussed as an example of successfuldevelopment of a neuroprosthetic implant.

Fig. 6. Summary of optimum surface behaviour of an

implanted neuroprosthesis: encapsulation surface (left),recording electrodes (left bottom) and stimulating electro-

des (right bottom).



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In a healthy ear, the sound vibrations are col-lected by the outer ear and sent down the ear canalto the eardrum which vibrates the three small bones

of the middle ear. Subsequently, the uid in thesnail-shaped cochlea vibrates, and these vibrationsstimulate the approx. 30,000 tiny hair cells whichare located in the organ of Corti inside the cochlea.The electrical signals produced by the hair cellsstimulate the bipolar cells of the spiral ganglion,which central bres form the auditory nerve whichleads the signals into the brain where they are inter-preted as sound. The sound is tonotopically rep-resented in the cochlea: high frequencies (up to 20kHz) are sensed in the basal region, whereas low fre-quencies (down to 16 Hz) are sensed in the highestpart of the cochlea, the apex. This distribution hasits origin in a 104:1 gradient of the stiness of thebasilar membrane and is the prerequisite of the

tonotopic organisation of the auditory system.Besides this site-dependent principle of frequencydiscrimination which is called spectral analysis,there is another mechanism called periodicity analy-sis. Here the frequency is ``calculated'' in the brainbased on the time period of incoming action poten-tials.

In case of a disease or a trauma, the hair cellsmay be diminished or damaged. The same may hap-pen at increasing age. A widespread reason foracquired hearing impairment or even deafness ismeningitis, where the site of damage is almostalways the cochlea with loss of the organ of Corti.As a consequence of hair cell loss, the auditorynerve may not be stimulated, and even the loudest

sounds may not be heard. That is why hearing aidswhich only amplify loudness of sound do not helpin inner ear diseases. When the auditory nerve itselfstill is intact, it may be stimulated articially by elec-trodes implanted in the cochlea, and this is what isdone with the so-called cochlea implant which isapplied now for approx. 20 years. The success variesdue to heterogeneity of diseases and the dierentdegree of destruction. When the auditory nerve itselfis destroyed in addition to hair cells, e.g. due to sur-gical removal of bilateral tumours of the acousticcanal, a cochlea implant does not make sense anymore. In such cases, direct stimulation of the brainareas may be tried in order to elicit acoustic sen-sations.

4.1. Brain Stem Implants

Such a stimulation can be performed by the audi-tory brain stem implant (ABI), which electricallystimulates the auditory pathway at the level of thecochlear nucleus. For this purpose, it could beshown that penetrating electrodes are more eectivethan surface electrodes, and their necessary stimu-lation threshold was considerably lower (67.5 vs11.4 mA), their dynamic range was bigger (13.1 vs24.5 dB), and metabolic activity of some CNSregions was higher as could be shown by higheruptake of [14C]2-deoxyglucose (El Kashlan, 1991).The correct placement of an ABI may be aided by

recording of brain stem responses which arerecorded during the implantation (Waring, 1995,1996). The evoked response generally had 2 or 3

waves, and peak latencies of these waves wereapprox. 0.3, 1.3, and 2.2 msec (Waring, 1995).

McCreery et al. (1992) also utilised evoked poten-

tials to optimise the position of stimulation micro-electrodes in the cochlear nucleus in experimentswith cats. They used 75 mm iridium wires for stimu-lation, and neurons and neurophil adjacent to themicroelectrodes appeared to remain undamagedexcept for some small gliotic scars. However, a cer-tain depression of neuronal excitability was foundwhich implies the necessity to nd the lowest poss-ible excitation threshold by careful monitoring ofboth the psychophysical threshold for auditory per-cepts and the electrically evoked auditory brain stemresponse.

Brackmann et al. (1993) described implantation ofa single-electrode ABI into the auditory brainstemof patients who were totally deaf due to removal of

their bilateral acoustic neuromas. Correct position-ing of the electrode in the lateral recess of the fourthventricle is very important for maximised auditorysensation and minimised activation of other nerves.Shannon et al. (1993) reported that these implantedelectrodes were stable for more than 10 years, andthat the auditory sensations produced by theimplant were similar to the results achieved bysingle-electrode cochlea implants. Patients could dis-criminate sound when it was combined with lip-reading. By this implant, tinnitus reduction could beachieved in several patients (Soussi and Otto, 1994).Later on, also implants with eight electrodes havebeen applied (Otto and Staller, 1995). Twelvepatients received such an implant, and eleven of

them received useful auditory sensations.

4.2. Cochlea Implants

Though there are slightly dierent views aboutwho is a candidate for cochlear implantation, it iswidely accepted that recipients of such an implantshould be at least 2 years old, have a severe or aprofound bilateral hearing loss and receive little orno benet from conventional hearing aids.Successful application of a cochlea implant requiresproper training, thus the recipients must have a highmotivation. An appropriately working device isexpected to allow the patient to detect speech andenvironmental sounds, to use the telephone, to

improve lip-reading abilities, to distinguish betweendierent kinds of environmental sounds and, in caseof children, to improve speech and language learn-ing.

How does the cochlear implant work? As outlinedbefore, it by-passes damaged parts of the ear anddirectly stimulates nerve bres of the auditory nerve.These signals may be interpreted in the brain as asound after a period of rehabilitation. The wholedevice consists of two parts, one situated outside thehead (microphone, speech processor, power supply,transmitter) and one implanted into the ear (recei-ver, stimulating electrodes). The sound recorded bythe microphone is converted in a series of electricalsignals by the speech processor which is then trans-

mitted through the skin to the receiver (Fig. 7). Thesignals are led to the electrodes, which apply the sig-nal to the auditory nerve bres. The right program-



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ming of the speech processor starts 46 weeks afterimplantation, and this may take several weeks tomonths. The whole procedure of programming istightly connected with adaptation processes in thebrain which may require synaptic plasticity in orderto rebuild the auditory system.

Although there are 30,00040,000 nerve bres inthe auditory nerve, stimulation can be performedfor technical reasons by only few electrodes.

Commercial devices usually have 622 electrodes.Based on the tonotopic organisation of the auditorysystem described above, it would be straightforwardto arrange stimulation microelectrodes along thewindings of the cochlea and apply distinct electricalstimulation pulses according to the occurrence ofdistinct frequencies in the environmental soundrecorded by the microphone. The speech processorwould analyse the recorded sound with respect todistinct frequency ranges and generate an appropri-ate series of electrical pulses for every single stimu-lation electrode. Indeed, this approach is performedwith a variety of commercial multi-channel systemsand there has been good success upon implantationof such devices. The number of stimulating electro-

des mentioned above should be increased in order toachieve a higher quality of stimulation and thus abetter acoustic sensation by the patient. However, if

the number of electrodes is increased, their distancegets smaller, and considerable interference andcross-talk between the channels occur (Hartmann

and Klinke, 1990). One way to resolve this problemand to minimise interference was development ofnew signal processing schemes where only one elec-trode is activated at a given time. This method iscalled ``sequential stimulation'' in dierence to sim-ultaneous stimulation. However, spatial resolutionof electrical stimulation cannot be enhanced as itwould be desirable.

As described already for the brain stem implants,also with cochlear implants measurement of evokedpotentials is used to evaluate eects of electricalstimulation. Whereas most information can begained by recording of responses of individual cellsor bres (Merzenich et al., 1973; Glass, 1983;Hartmann et al., 1984; van den Honert and

Stypulkowski, 1984), measurement of evoked poten-tials is much more convenient, can be performed inalert animals and can be used for the investigationof long-term changes. Among the evoked potentials,recording of the middle latency responses seems tobe useful (Kileny and Kemink, 1987; Burton et al.,1989), and extended investigations have been per-formed, e.g. by Popela r et al. (1993, 1995). Theauditory brainstem response (ABR) is often used toevaluate the residual hearing of patients as well asthe performance of cochlear implants. In most clini-cal cases, click-evoked ABRs are recorded. Ofcourse, measurement of evoked potentials is not ahearing test per se.

In the rst implants, only a single electrode has

been used, and provided therefore limited success.Cochlear implants gained higher acceptance sinceMEAs have been applied which allow multi-channelstimulation. The electrodes are arranged on a whorl-shaped carrier which is introduced into the cochleaso that the single stimulating sites are placed near tothe appropriate sites of the organ of Corti and thesurface of the cochlear nucleus. However, it must betaken into account that damage of components ofthe cochlea can occur by the surgical process ofinsertion of the long electrodes applied for multielec-trode stimulation. The spiral ligament is rst site ofdamage in these cases, and also the basilar mem-brane could be disrupted which would lead to alesion of remaining dendrites and subsequent de-

generation of the spiral ganglion cells. This meansthat potentially existing residual hearing may bedestroyed. On the contrary, there are reports thatinserted implants caused minimal damage and arewell tolerated (Burton et al., 1996). In order to over-come the problem of cochlea damage, several ``soft''surgical strategies have been developed (e.g. use oflubricating liquids for better insertion, Rogowski etal., 1995) which will not be further discussed here.Cohen (1997) critically contemplates the concept of``soft surgery'' and states that success of the cochleaimplant mainly depends on full electrode insertion,the right stimulation strategy and the survival of asucient number of ganglion cells.

The problem of full electrode insertion mentioned

above arises particularly in the case of cochlea ossi-cation which often happens. The implant cannot beinserted into the cochlea any more. Drilling is only a

Fig. 7. (A) Scheme of a cochlea implant with arrows indi-cating informatiom ow. Sound is recorded by a micro-phone and converted into a series of electrical signals bythe speech processor. The electrical signals are transmittedthrough the skin to the receiver which converts them intostimulating pulses applied by the stimulator to the neuronsof the auditory nerve. 1ear canal, 2eardrum mem-brane, 3three little bones (hammer, anvil, stirrup), 4cochlea, 5auditory nerve. (B) Detailed view of thecochlea with the inserted stimulator. The electrodes are dis-

tributed along the implant thus providing spatially dieren-tiated stimulation. 1scala tympani, 2scala vestibuli, 3basilar membrane with the organ of Corti which containshair cells and sensory nerve endings, 4ganglion spirale,

5auditory nerve.



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partial solution because the drilled out portion islimited to the rst part of the basal turn, and stimu-lation results are poor. One possibility is not to

insert the multi-electrode carrier into the cochlea atall, but to drill small holes into the cochlea at cer-tain sites and insert small single electrodes throughthese holes (Chouard, 1994). Although this pro-cedure appears to be delicate and time-consuming, itlooks more advantageous, particularly because pos-itions of the electrodes may be set very exactlywhich favours tonotopic excitation. Another way isthe development of an implant consisting of twoarrays (Lenarz et al., 1997). Whereas the rst arrayis inserted into the drilled out basal turn, the secondarray is inserted into the second turn which has tobe opened for this purpose. The authors reportedthat an improved performance of the implant couldbe achieved.

In this context, it was argued that the tonotopictheory for cochlear implants is not valid becausedendrites of the spiral ganglion cells may retract ordegenerate after loss of the hair cells (which is dis-cussed later), and therefore multi-electrode implantswould not be useful. It thus may be concluded thatselective stimulation of the spiral ganglion cellswould not be possible, and consequently only oneelectrode with an electrical eld penetrating thewhole cochlea would be enough. However, it is wellestablished that a tonotopic organisation occursboth in the healthy auditory system and in the caseof deafness. Naturally, quality of the tonotopystrongly depends on the time of onset and durationof deafness. In any case, the performance of patients

with multichannel cochlear implants is anticipatedto be maybe better than the performance of patientswith single-channel devices (Gantz et al., 1987;Tyler, 1987). Experiments with deafened catsshowed variations in parameters of auditory evokedpotentials recorded in individual tonotopical corticalplaces when the auditory nerve was stimulated withdierent congurations of electrodes through amulti-electrode implant (Popela r et al., 1995).Stimulation with monopolar electrodes yieldedresults with much greater variability among individ-ual cats and induced profound functional alterationsin the CNS (Leake et al., 1995). For this reason, theauthors stated that application of monopolardevices should be contra-indicated for at least young

children. Comparative studies in humans showedthat the outcome of an implant is better with ahigher number of active electrodes (Amadori et al.,1996).

In 1995, more than 12,000 people world-wideused cochlear implants (for an overview, see, forexample, NIH Statement, 1995). A majority of themis able to understand up to 80% of high-context sen-tences, recognises environmental sounds and can lis-ten to music. In spite of good experience with thesedevices, there are also limitations. Noisy environ-ment remains a problem. Cochlear implants maynot provide the dynamics of sound, i.e. the range ofamplitude, like normal hearing does. While normalhearing provides loudness dierences of 4 orders of

magnitude, only a factor of 10 may be heard bysuch an implant. Another problem is that success ofimplantation and rehabilitation cannot be predicted

very exactly. Several deaf people still cannot use thetelephone and depend on lip-reading, and some ofthem are helped only slightly by these implants.

Failure to surgically recover auditory function isparticularly observed when individuals were bornwith deafness. Poorer results were also achieved inchildren with pre- or peri-lingual onset of deafnesscompared to children with post-lingual onset ofdeafness. However, the dierence between these twogroups appears to lessen with time (NIH Statement,1995). The reason for this failure is most probablyabsence of the appropriate synaptic connections inthe brain which would allow to interpret the incom-ing nerve signals as sound. For acoustic sensations,especially with patients who are deaf from birth, alot of connections have to be established in thebrain. This task is certainly easier to accomplish in adeveloping brain than in an adult one. This may

explain the observation that implantation at an ageof 2 years ultimately results in a better auditory per-formance than implantation at the age of 3 years orlater (NIH Statement, 1995). Special problems ofthe application of cochlea implants in children arereviewed by Langman et al. (1996).

The observation that individuals with shorterauditory deprivation achieve better results thanpeople with a longer deaf period can be seen in thesame context as the ability of the brain to processsignals incoming from the auditory nerve. It hasbeen known for a longer time, that sensory depri-vation leads to plastic changes in the correspondingareas of central nervous system, particularly in thecortex. In case of removal of sensory input, the

deprived area of the somatosensory cortex becomesresponsive to neighbouring regions (Kaas et al.,1983; Kaas, 1991).

Another problem is that often not only hair cellsbut also neurons of the spiral ganglia are aected bythe disease or trauma which lowers the prospects ofa good auditory performance of a cochlea implant.There are hints that hair cells in the vestibule maybe regenerated (Forge et al., 1993; Warchol et al.,1993), but these ndings are controversial (Rubel etal., 1995). Moreover, hair cells in the cochlea aremuch more dierentiated, and regeneration of losthair cells in the mammalian cochlea appears not tobe possible at the moment. First eect after loss ofhair cells is retraction and degeneration of the den-

drites of the bipolar spiral ganglion cells. It is gener-ally known about the nervous system that neuritesmay retract when the target cells are lost becauseneurotransmitters and other factors are not suppliedany longer. In the cochlea, BNDF and NT-3 areproduced by the developing organ of Corti, andNT-3 is produced by the inner hair cells also in theadult stage (Ylikosi et al., 1993; Schecterson andBothwell, 1994). Moreover, transcripts for trkB andtrkC which are specic receptors for BDNF andNT-3, respectively, were found on the auditoryneurones (Ylikosi et al., 1993). Recently it wasfound that NT-3 can elicit a tropic response on out-growing auditory neuronal processes in vitro(Malgrange et al., 1996). In the guinea pig cochlea,

auditory neurones underwent apoptotic cell deathafter destruction of associated hair cells, and per-fusion of BDNF and/or NT-3 onto the scala tym-



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the inuence of the stimulating electric elds. Thus,excitation of these axons markedly lowers localspecicity of stimulation at the site of implantation.With a subretinal implant, the electrodes could per-form directly the function of photoreceptor cells bygiving a stimulus upon irradiation. Subsequent in-formation processing would be performed by theother neuron layers of the retina. This would not bepossible with an epiretinal implant, and thereforethe latter would require a highly sophisticated

encoding of visual information into stimulating sig-nals.

Numerous experiments have been performed toachieve excitation of RGC cell bodies rather thanaxons. This task is complicated because the detailedposition of stimulating electrodes relative to the cellbodies cannot be controlled. One possibility is toutilise dierent excitation thresholds of cell bodiesand axons upon cathodic or anodic stimulation, andit seems that preferred excitation of cell bodies canbe achieved by anodic stimulation. Anotherapproach is application of non-radial electrode con-gurations. When the lines of the electrical eld oflinear electrodes arranged in parallel run perpendi-cularly to the axons, a minimal potential dierence

is built up within the axon, and in the other case,with the electrical eld longitudinal with the axon,the potential dierence would be higher thus lower-

ing threshold for excitation. Currently, geometriesof MEAs are under investigation with non-sym-metrical electrodes in order to improve selectivity ofganglion cell stimulation.

With subretinal implants, stimulation is supposedto be more similar to the natural way becauseincoming light would be applied as an electrical sig-nal directly at the place where it hits the retina, andthe same is done by the photoreceptor cells.Consequently, the concept of most subretinal

implants consists of an array of photodiodes whichdirectly supply an electrical pulse upon light ex-posure. Further processing of these signals wouldthen be carried out by the neuronal layers of the ret-ina, and RGCs would nally transmit the visual in-formation to the brain. Naturally, neuronal layersbeneath the ganglion cell layer are necessary for thisconcept. However, it must be expected that theydegenerate step by step because photoreceptorswhich supply stimulating signals are lost. In fact,such a degeneration has been observed in all retinallayers, but it is claimed that remaining neuronsshould be sucient for the application of a visualprosthesis (Santos et al., 1997). Retinal implants areinserted through the anterior part of the eye, and

animal experiments are performed mainly with rab-bits. The whole surgery is very delicate because theretina is very thin and soft. For epiretinal implants,

Fig. 8. Examples of investigated possibilities to (partial) restoration of vision by the application ofmicroelectrode arrays (MEAs). If the retinal ganglion cells (RGC) are still intact, then the stimulating

MEA could be implanted into the eye, either onto the RGC layer (epiretinal implant) or by replacingphotoreceptor cells (subretinal implant), and nerve signals of stimulated RGC are forwarded into thebrain. Otherwise, the MEA is put on the surface of the visual cortex (epicortical implant). Alternatively,

the cortical implant can also consist of an array of needles which are inserted into the cortex.



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lation by many electrodes. Such a matrix was simu-lated by a monitor covered with an opaque perfo-rated mask, and from the experiments was

concluded that a MEA consisting of 25

25 electro-des on an area of 11 cm2 should produce a phos-phene image on a visual acuity of approximately 20/30 provided that the MEA is implanted near thefoveal representation of the visual cortex (Cha et al.,1992).

Schmidt et al. (1996) implanted 38 microelec-trodes in the right visual cortex of a 42-year-oldwoman who had been blind over 22 years. 34 micro-electrodes were able to produce phosphenes, andmost of the microelectrodes had stimulationthresholds below 25 mA. Brightness and size of thephosphenes could be inuenced by stimulation par-ameters, and separate phosphenes could be detectedby the patient when stimulating electrodes had a

spacing of 500 mm. Moreover, the authors reportedthat six phosphenes could be elicited simultaneously,and they

Implantable Bio Electronic Interfaces for Lost Nerve Functions

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