Implantable bioelectronic interfaces for lost nerve functions

Progress in Neurobiology Vol. 55, pp. 433 to 461, 1998 # 1998 Elsevier Science Ltd. All rights reserved Printed in Great Britain 0301-0082/98/$19.00

PII: S0301-0082(98)00013-6

IMPLANTABLE BIOELECTRONIC INTERFACES FOR LOST NERVE FUNCTIONS P. HEIDUSCHKA* and S. THANOS*$ *University Eye Hospital MuÈnster, Experimental Ophthalmology, Domagkstraûe 15, D-48149 MuÈnster, Germany (Received 6 January 1998) AbstractÐNeuronal cells are unique within the organism. In addition to forming long-distance connections with other nerve cells and non-neuronal targets, they lose the ability to regenerate their neurites and 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 arti®cial implants prompted to create particular devices 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 applicability and functional perspectives. (1) Successful implants like the cochlea implant and peripherally implantable stimulators are discussed. (2) Less developed and not yet applicable devices like retinal or cortical implants are introduced, with particular emphasis given to the reasons for their failure to replace very complex functions like vision. (3) Material research is presented both from the technological aspect and from 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. The review ends by pointing to future perspectives in neuroimplantation and restoration of interrupted neuronal pathways. # 1998 Elsevier Science Ltd. All rights reserved

CONTENTS 1. Introduction 2. Implantable electrodes 2.1. General remarks 2.2. Electrodes implanted without nerve cut 2.2.1. Cu€ electrodes 2.2.2. Penetrating electrodes 2.3. Electrodes for regenerating nerves 3. Biocompatibility 3.1. General remarks 3.2. Implant materials 3.3. Response to implantation 3.4. Surface modi®cation 3.4.1. Surface roughening 3.4.2. Chemical modi®cation 3.4.3. Speci®c modi®cations 4. Auditory implants 4.1. Brain stem implants 4.2. Cochlea implants 5. Visual implants 5.1. Retinal prostheses 5.2. Cortical prostheses 6. Other neuroprosthetic implants 6.1. Bladder stimulation 6.2. Stimulation of spinal cord and brain 7. Final conclusions References

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$ Author for correspondence. Tel.: 49 251 83 56915; Fax: 49 251 83 56916; e-mail: [email protected]. 433

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ABBREVIATIONS ABI ANN BDNF CNS ECM FES IFN IL

Auditory brain stem implant Arti®cial neural network Brain-derived neurotrophic factor Central nervous system Extracellular matrix Functional electrical stimulation Interferon Interleukin

1. INTRODUCTION Diseases and accidents associated with damage of nerves, in particular within the CNS, often have dramatic consequences. The ®rst reason is that appropriate regeneration of the central neuronal connections and restoration of synaptic connections are not possible in most cases. This results in failure of functional recovery. The second reason is that dying and disposed neurons cannot be replaced by new neurons. This cascade of interactive events results in glial proliferation and in inadequate repair, called gliosis. In contrast to CNS, bundles of the peripheral nerve system like in the limbs display di€erent responses to experimental or accidental cut: the axons can regenerate, the cell bodies do not degenerate and there is virtually no need for gliosis and replacement of the lost neurons. De®cits along peripheral nerves may be reconstructed spontaneously or surgically, and physiological properties of the function are regained after regeneration of axons and synapse formation. The ®nal functional use of the corresponding target deserves some training, as directly associated with synaptic reorganisation, potentiation and trophic in¯uences at the neuromuscular endplates. However, such restorative events are limited to less complex and unilaterally oriented pathways, like the innervation of a single muscle or the arrangement of nerves with topologically and functionally distinctive targets. In spite of the ability of peripheral nerves to regenerate, surgical repair of nerve damage is not as easy or even not possible for several reasons: . the nerve is not accessible to surgically join the proximal 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. Diculties in surgery apply especially to the CNS where most surgical interventions are not possible without destroying neighbouring parts of the nerve tissue, where the network-like highest complexity of multiple circuits has been formed during development. In this network, molecular guidance does not properly work any more, and oligodendrocytes seem to form an environment which inhibits axonal growth. Moreover, damaged neurons can be destroyed by additional local mechanisms, probably developed to eciently remove the sick elements and preserve the remaining structural and functional

MEA NT-3 PET PTFE RGC RP TNF VEP

Microelectrode array Neurotrophin-3 Polyethylene terephthalate Polytetra¯uoroethylene Retinal ganglion cells Retinis pigmentosa Tumour necrosis factor Visually evoked potential.

integrity of the tissue. Finally, astrocytes are responsible to ®ll the structural gaps with proliferation and to communicate will all other elements, thus balancing the de®cits. Frequently, whole organs or part of the body are removed or destroyed, e.g. a limb is lost which cannot be replaced naturally, ultimately demanding an external intervention. In such cases, rudimentary organ of body functions (stability, standing up, walking) can be in part compensated with external mechanical prosthetics. However, they are not acting on the neural function, but replace the peripheral muscular, bone or joint functions. It seems desirable, but only in extremely rare cases possible to date, to develop prosthetic interfaces between the nervous system and its various peripheral targets. These diculties are multiple and require ®rst the profound understanding of the neuronal circuitry and function and second a directed ¯ow of neuronal information from its natural origin through a sensing prostheses into a ®nal target in a functionally remodelled form. A classical injury resulting in devastating functional impairments is the incision of spinal cord with usually combined lesions of ascending and descending paths, in addition to local necrosis of intraspinal neurons. One of the predominant ®elds of prosthetic intervention in the future will be the cure of paraplegic and paralytic individuals with best results until now towards to reconstruct the simpler re¯exes like that of the urinary bladder function. Other work is carried out in order to stimulate muscles of the legs and the back in order to make possible getting up, standing or even walking. First attempts to achieve regeneration of neural connections through regenerating nerves in the spinal cord have been made, but they are still far away from restorative applicability. Smaller, well integrated sensory organs and compartments may be more accessible to so-called arti®cial sensors where a direct connection between electronic devices and nerves is needed. The probably most popular application where di€erent devices work bene®cially is the cochlear implant. Conceptually, further sensory organs may be replaced 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 or epiretinal implants. In addition, further approaches were devoted to replacing the optic nerve with peripheral nerve implants, with limited success in regaining function, but encouraging results towards understanding the mechanisms of axonal regener-

Implantable Bioelectronic Interfaces for Lost Nerve Functions

ation and opening a new ®eld of natural neuoprosthetics in combination with arti®cial prosthetic devices. It is desirable, however, that all arti®cial approaches are accompanied by endeavours to facilitate the natural approach, i.e. the regeneration of axons towards their natural targets and fresh formation of new synapses leading to at least partial functional recovery. The goal of this work is to critically review some aspects of these approaches, particularly the required functional abilities of implants, the design of microelectrode arrays (MEAs) for optimal function and bio-(neuro-)compatibility, their biologically proper implantation and their function for recording of nerve signals and/or stimulation of nerves and/or muscles.

2. IMPLANTABLE ELECTRODES 2.1. General Remarks The nervous system is the most complex biological system and is therefore characterised by high vulnerability to external violence and genetic disturbances. Profound knowledge about its development, functional consolidation and structural organisation is prerequisite for any attempt to establish good recording or stimulation in order to treat defects in a proper way. A number of proposed ideas for electronic implants is based on a more ``engineer-like'' way of thinking than considering the complexity of the biological systems. For instance, nerve ®bres are not ``wires'' or ``cables'' in the mechanical point of view, but long, sensitive structures consisting of biological membranes with multiple receptors and delicate interactive elements for online sensing the environment and transmitting information via molecules, potentials and changes in the chemo-electrical activity. The principal requirements to any implantable structures are therefore features mimicking some of the biological functions of nerves and replacing these functions depending on the scope of implantation. Many e€orts have been undertaken to interface electronics to nerves, and there is a reasonable number of successful applications with di€erent goals. The aim of using implants in basic research is to understand working principles of the brain, to analyse processing of information, to unravel principles of connectivity within the nervous system and to study cell±cell interactions in subsets of neuronal populations. Another goal is to replace functions which are lost due to damage of the nervous system. In this context, some attempts are concentrated on sensory functions and stimulation of muscles by implanted electrodes and on development of closedloop ``neuronal prostheses''. Microtechnology and microelectronics have obtained impressing capabilities which have to be combined with possibilities of modern biology and medicine under careful consideration of possibilities and constraints of neuroanatomy and neurophysiology. In fact, implantation of electrodes for nerve signal recording and nerve stimulation has been carried out for some couples of

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years with success, and there are also some promising concepts for the future and interesting approaches for special applications. A review about the basics and di€erent aspects of neural prostheses is given by Agnew and McCreery (1990a). The success of neural prostheses depends basically on their capability to record nerve signals and/or to stimulate nerves and muscles. It is obvious that such implants must ful®l very special requirements which are directed to the electrodes and the substrate which carries the electrodes. Prerequisite of attachment to nerves and sustaining there is the biocompatibility for nerve-speci®c implants, or more strictly spoken, their neurocompatibility as shall be addressed later. The substrate which carries the electrodes must be completely insulating to prevent cross-talking between them. The choice of the material will also depend on the possibilities and restrictions arising from the fabrication process. In microelectronics, silicon is most frequently used. However, in the majority of implantation cases, it will be of favour to use ¯exible material for the implants in order to mimic biological tissue and to reduce the possibility of mechanical damage. As one of these ¯exible materials, polyimide obtained attention in the last years. As they are directly in contact with the nerve of interest, the appropriate construction of the implanted electrodes is of key importance for the success of the whole arti®cial system. They have to be of a size comparable with the size of the neurons in order to interact only with few or even one neuron. If electrodes are intended to interact with nerve ®bres, i.e. axons, in the case of myelination their ``radius of action'' has to be sucient to reach the nearest node of Ranvier. The small size requires optimised geometry and electrical properties for suf®cient recording ability and a high relative charge transfer capacity. Moreover, all components have to be adapted to this, i.e. electronic units are needed with a high sensitivity, a high signal-to-noise ratio, and sucient shielding between single channels within the device and against external interferences. 2.2. Electrodes Implanted Without Nerve Cut 2.2.1. Cu€ Electrodes The term ``cu€ electrodes'' applies to those devices which engulf the entire circumference of a nerve. The shape of these electrodes or of the array of electrodes has to be adapted to the natural arrangement and thickness of neurons or axons by also taking in account the vascularisation at the site of implantation. Cu€ electrodes are applied preferably at peripheral nerves. The advantage of cu€ electrodes is that implantation is relatively easy and that the nerve is not damaged by proper surgical implantation. Possible displacement along the nerve bundle can be circumvented by mechanical ®xation at the site of interest. First models carried only one or two electrodes. They were made using a platinum foil (``split-cylinder'', Avery, 1973) or platinum wires embedded into insulating material (``wrap-around'',

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Hagfors, 1972). The ®rst models have been rather sti€, often resulting in damage of nerves due to mechanical displacement, or pressing the nerve ®bres, or disturbing the vascularisation and causing ischemia. Modern cu€ electrodes try to avoid this by the introduction of ¯exible materials and adaptable geometries, like a helix-shaped electrode (Agnew et al., 1988) or a spiral-cu€ electrode (Rozman, 1991) which allow adjustment of the implant to the actual diameter of the nerve ®bre bundle. Another possibility is the application of socalled ``half-cu€'' electrodes ®rstly described by Kim et al. (1983) and patented by Testerman and Bierbaum (1994). A particular design is the arrangement of ¯exible interdigitating sub-units with microelectrodes along a backbone-like carrier (Klepinski, 1994; Meyer et al., 1995). Models of these two designs are shown in Fig. 1. First models of cu€ electrodes did not completely ful®l the requirements for accurate measurements, because they only allowed recording of super®cial sum potentials with the axons in the centre of the nerve cylinder to contribute less signi®cantly to the measured signal. In accordance, the inner axons were less a€ected by stimulation than the super®cial axons of the bundle. For this reason, advanced models of cu€ electrodes have been developed with more smaller electrodes arranged around the per-

imeter of the ®bre bundle which allowed to approach central axons too. Topographic stimulation of nerve ®bres within a given nerve bundle may sometimes be important, because di€erent muscles may be innervated by the ®bres of selective localisation within the bundle, and often a mixture of e€erent and a€erent ®bres occurs in the nerve. Such mixed populations may lead to undesirable sensations during stimulation. In the case of several nerve ®bres innervating di€erent muscle ®bres of one muscle, cyclic stimulation can be performed which guarantees overall stimulation frequency, while stimulation frequency and thus fatigue of a single muscle ®bre can be low (Happak et al., 1989; Talonen et al., 1990). With a sucient number of electrodes within the cu€, high selectivity of stimulation can be achieved by either longitudinal or transversal currents produced by appropriately switched electrodes (Veraart et al., 1993). However, there is still the problem of possibly di€erent ®bre diameters which leads to di€erent activation thresholds of the ®bres. Goodall et al. (1996) found that large ®bres were activated before smaller with a cu€ electrode containing 12 electrodes arranged in four longitudinal tripoles, irrespective of ®bre position. Position selectivity could be enhanced by a higher 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€'' electrode 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 ¯exible 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).

Implantable Bioelectronic Interfaces for Lost Nerve Functions

437

application of such sophisticated electrode designs and stimulation protocols is detailed knowledge of the distribution of nerve ®bres within the nerve bundle and their function. Moreover, a really reliable ®xation of the cu€ must be achieved because otherwise the cu€ may rotate around the nerve or shift along the nerve, both leading to loss of selectivity. 2.2.2. Penetrating Electrodes The so-called penetrating electrodes are another type of electrodes aiming to directly approach nerve ®bres situated deeper in the tissue. The ®rst penetrating electrodes were simply thin metal wires or needles which were inserted into the nervous tissue. They often were insulated except a region on the tip. The other possibility was to use glass micropipettes pulled out to very small inner diameters and ®lled with saline. They also can be inserted into nervous tissue, and signals are recorded by a thin metal wire electrode within the pipette. For simplicity, but wrongly, the glass micropipettes themselves often are designated as ``electrodes''. One recent example for the use of such pipettes is given in the work by Welsh et al. (1995), who investigated the role of the inferior olive, a major cerebellar a€erent, by recording the activity of the Purkinje cells. For this purpose, they positioned up to 39 glass micropipettes with electrodes (®lled with saline, 1±2 MO, 2±4 mm tip diameter) independently 100±125 mm below 3 mm2 of the cortical surface of rats, with a distance between the electrodes of 250 mm. This procedure is really time-consuming, and its application is limited by its complexity. The animals have to be anaesthetised and held in a ®xed position. Moreover, the position of the electrodes is not optimised with respect to the location of the desired target, in this work the Purkinje cell layer, but recorded signals are used from those electrodes which are placed at best by chance. Wedge-shaped microprobes carry a line of electrode sites for recording and stimulation and may therefore 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 the range of 2 mm for the shank length, 100 mm for the width and 20 mm for the thickness, but there are already smaller structures in development. An onchip electronic circuitry would allow pre-ampli®cation and ®rst processing, but it is not yet realised in most cases. Howard et al. (1996) built a recording array where they combined high-impedance microelectrodes with low-impedance EEG electrodes, and activity of human cortical neurons could be recorded. For signal recording and stimulation over a larger area, e.g. areas of the cortex, two-dimensional MEAs are necessary as they are known as solid planar arrays for in vitro experiments (Wilson et al., 1994; Nisch et al., 1994; Gross et al., 1995; Bove et al., 1997). A ¯exible planar MEA with polyimide as substrate and 24 gold microelectrodes (40  40 mm, 210 mm spacing) was made for recordings of electrical activity of the cortex (Owens et al., 1995). Improvement of recording structures was achieved by silicon microtechnology which made it possible

Fig. 2. Di€erent layouts of penetrating electrodes. (A) Wedge- or shank-shaped electrode carriers are one of the most common designs developed and applied by several groups. In many cases, more than one shank are combined. Typically, the shanks are 1±3 mm long, 30±100 mm wide and 8±20 mm thick. (B) Proposed structure of an array of needles carrying multiple electrode sites thus reaching a really three-dimensional arrangement. (C) A combination of cu€ and penetrating electrodes (Durand and Tyler, 1995; Tyler and Durand, 1997). Only one possible design of the implant is shown in opened and closed states. The electrodes penetrate the nerve when the cu€ is bent around the 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.

to create not only planar two-dimensional MEAs but also penetrating electrode structures for in vivo measurements. The substrate carrying the electrodes is either needle- or wedge-shaped to allow penetration of the nervous tissue which makes possible recording from and stimulation of axons not only on the surface but also in a well-de®ned depth within the tissue, e.g. within the ®bre bundle or regions of the brain. Implanting such a device is also associated with an accidental damage of the tissue, i.e. some neurons will be destroyed, and a certain portion of the axons will be disrupted. Moreover, sti€ness of many models may lead to damage of nervous tissue. That is why the e€orts are directed to miniaturise the penetrating parts of the implant and to use more ¯exible materials. Nordhausen et al. (1996) reported about a siliconbased two-dimensional MEA shaped as a grid of 10  10 needle electrodes with a spacing of 400 mm. The needles are approx. 80 mm wide at the base and 1.5 mm in length. They are insulated with polyimide except for approx. 50 mm at the tip, the latter being coated with platinum to form the active electrode. The electrode array was successfully applied for the recording of local visually evoked responses in the visual cortex of cats at sub-sets of 15 electrodes (Nordhausen et al., 1996), and these measurements were continued in order to re®ne recording and signal processing procedures (Maynard et al., 1997). A similar approach was performed by Rutten et al. (1995). They created a three-dimensional needle array with 128 recording sites with one electrode on the tip of a needle. The needles are made from silicon and are embedded into a glass substrate. They vary in height from 250 to 600 mm and have a distance of 120 mm, with a tip size of 15  15 mm. The di€erent length of the needles should allow reaching

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approximately three-dimensional areas of nervous tissue, e.g. in a peripheral nerve fascicle. It could be suggested that the optimal MEA would be one with the possibility to move each single electrode separately to its best place in order to optimise recording or stimulation. On the one hand, this is limited by the possibilities of microtechnology, because it would be very complicated to design and fabricate a MEA with independently movable and mechanically stable electrodes with perfect conduction of signals and insulation in the aqueous surrounding. On the other hand, there is the biological issue to ®nd out which place within the nervous tissue is actually the best for the desired purpose. This is time consuming, and it is dicult to predict how much of the nervous tissue is damaged during penetration of the electrodes and their lateral movement. In view of these problems, it would be useful to combine the approaches mentioned above and create a three-dimensional MEA by arranging needle probes in a grid with electrodes placed along the probe shank as shown in Fig. 2(B). This would provide a high-density three-dimensional arrangement of electrode sites, and best of them could then be found out after successful implantation by testing recording or stimulation characteristics. Of course, there is a huge number of technical problems associated with the very small dimensions of MEAs and also of the design shown in Fig. 2(B). One of the problems is that, on the insulating substrate, conducting electrode sites and leads must be placed. A second problem is that cross-talk between the leads and leakage have to be avoided. A third drawback is that with increasing number of electrodes the number of leads increases too, which demands an intricate on-chip design and reliable connections and cables to electronic processing units. Furthermore, the whole set-up must operate reliably over a long time period. Last but not least, processing of a high number of channels requires a sophisticated microelectronics which may not be too big and energy-consuming for practical every-day use. With the technological advance in the ®eld of constructing ¯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 possibly other elements of a microelectronic circuitry. The ¯exible nerve plate could be applied in a big variety of tissues and organs. For instance, it could be inserted into the retina or longitudinally between the ®bres of a nerve bundle or a muscle, and act therefore as a kind of penetrating electrode in these cases, or it could be placed onto the surface of special regions of the cortex depending of the scope of the experiment. The ¯exibility of the nerve plate would allow better protection of nerve tissue and very short distances between electrodes and neurons or ®bres which are necessary for an optimum decoding of nerve signals or transmitting stimuli to the nerves. As for cu€ electrodes, the geometry and the arrangement of penetrating electrode array variations are currently optimised to ®t with geometrical

nerve ®bre characteristics (®bre position in a bundle, diameter of the ®bres and their orientation). In¯uence of speci®c anode±cathode combinations and stimulus parameters are also under research. There are some research projects dealing with the development of a computer-assisted control of stimulating potential pulses for the formation of spatially distinguished electric ®eld within the nervous tissue which allow selective stimulation of the desired neurons. By sophisticated control of the height and time of applied potential pulses, a more discrete stimulation of the axons can be achieved. Such e€orts are undertaken for electrode arrays for both central and peripheral nerve systems. 2.3. Electrodes for Regenerating Nerves A di€erent approach is performed by the so-called ``electrodes for regenerating nerves''. They are designed as a MEA placed on a sieve-shaped (i.e. perforated) plate which contains holes which can be round or rectangular or even shaped as long narrow slots. The microelectrodes are situated nearby the holes or are part of the hole's wall in order to optimally record or to stimulate. The principal idea can be described very brie¯y: The nerve is cut ®rstly, then the electrode array is adapted into the expected path of the regenerating ®bres in a fashion that the nerve ®bres are allowed to regenerate through the perforations of the device (Fig. 3). The distal stump of the cut nerve is aligned at the opposite side of the electrode array in order to be used by the axons exiting from the perforations as a guidance path for further growth. In most cases, the perineural sheath of the nerve can be replaced in the area of insertion with polymeric tubes which act as mechanical stabilisers. The advantage of this approach is that with this device the electrodes are in intimate contact with the nerve ®bres, this allowing both accurate recording and ecient stimulation. Both procedures are expected to be performed relatively reliable because with a proper choice of the hole diameter a predictable number of axons would regenerate through individual perforations. Moreover, the microelectrodes remain always in the same position rela-

Fig. 3. Principal concept of regenerative electrodes. Axons regenerating from the proximal stump of a dissected nerve grow through a permissible array of electrodes carried by a substrate. The array may be designed as a kind of mesh or sieve. Latticed arrangements are also possible. The apertures (holes) of the substrate determine and ®x the position of the regenerating axons relative to the electrodes.

Implantable Bioelectronic Interfaces for Lost Nerve Functions

tively to the nerve ®bres because the ®bres are ®xed by the holes they grow through. The obvious drawback of this method is that the nerve has to be cut in order to regenerate through the implanted device. The success of the whole operation can be evaluated only several weeks later, when the axons have regenerated through the device. The second disadvantage is its applicability only in peripheral nerves, because central nerve pathways do not regenerate spontaneously. The third disadvantage is that the device limits the regeneration of some neurites, namely of those, whose growth cones hit on the device and fail to elongate within one of the pre-drilled holes. In spite of these limitations, such electrodes display an elegant way of application in the peripheral nerves. Indeed, attempts to develop and utilise such electrodes have been performed continuously since the sixties (for an overview, see Kovacs and Rosen, 1992). Mannard et al. (1974) combined 10 silver wires in a conical bundle. They were carried by a ¯attened epoxy bulb where holes have been drilled in. LlinaÂs et al. (1973) proposed a relatively advanced concept for an electrode device with a radial array of gold electrodes surrounding the holes. Loeb et al. (1977) reported on the fabrication of an electrode array with 0.3±1.2 mm long tubes the regenerating axons should be growing through. Edell (1980) developed a structure with long narrow slots for nerve ®bre regeneration. Electrodes were placed on the thin interspaces between the slots. Rosen and Grosser (1986) also proposed a concept of ``regenerative electrodes'' in order to ``restore normal nerve impulse communication and hence nerve function''. A micromachined silicon electrode was made by Akin and Naja® (1991). After implantation, nerve regeneration through this device could be achieved with the glossopharyngeal nerve of rats as model system (Bradley et al., 1992), and nerve signals could be recorded (Akin et al., 1994), even over a longer time period (Bradley et al., 1997). A complete set-up of a perforated electrode with a MEA was also developed and applied by Kovacs (1991). After implantation into the peroneal nerve in the hind limb of rats, some spontaneous action potentials could be recorded (Kovacs et al., 1992). Later on, more extended recordings from the rat peroneal nerve and the bullfrog cranial nerve were reported (Kovacs et al., 1994). The two last devices are shown in Fig. 4. Recently, a perforated MEA was described which was implanted between the cut ends of the rat sciatic nerve (Navarro et al., 1996). Another description of this MEA was given by Dario et al. (1997). The biological and technical aspects concerning the electrodes with perforating paths are complex and do not permit a simpli®cation in use, neither a generalisation in terms of their applicability, mainly due to the traumatic procedure of implantation. Although techniques of microfabrication have been developed rapidly during the last years, some problems arise if a large number of microelectrodes with the accompanying connections has to be placed and ®xed on a silicon chip. These problems are enhanced in electrodes for regeneration, because the holes take a big part of the total surface area and

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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.

limit per se the number of connections with individual holes. The shortage of surface of a microchip limits both the number and the size of holes for a given electrode. The minimum requirement for a perforating electrode is that growth cones ®t within the individual hole and pass through its lumen without disturbances in their growth properties, and later on, in their conduction velocities. It does not seem critical whether the holes are round or rectangular. Experiments of di€erent groups showed that, depending on the nerve where the chip is placed, the minimum size of a hole should be around 2±5 mm, and thus in the range of the thickness of a single axon. Some groups use hole sizes between 25 and 50 mm and expect that more than one axon can grow through each hole. The natural intention of a quantitatively high performance of regeneration requires 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 the microchip. A compromising way to use fewer connections with more bioelectronic interfaces is the multiplexing of the electrodes, although this process cannot be forced endlessly. Thus, the scope of the experimental design in individual cases determines the ratio between number of perforations and number of electrodes, again depending on the cross-sectional area of the nerve to be used. As an example, a ratio of 2.5:1 between the area of the chip bearing the holes and the cross-sectional area of the nerve is

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considered to be good for the regeneration of a big portion of axons (X. Navarro, personal communication). The physical need of space for the connections also determines a minimum possible distance between neighbouring holes. Nowadays, connection widths of only some mm can be realised, so that distances between the edges of holes of 10±20 mm are possible. A further critical point that is associated with the biological constraint of nutrition is the distance the regenerating axons have to get through the holes. In technical terms, this corresponds to the thickness of the chip. There have been some extreme cases, e.g. by Loeb et al. (1977), where that distance was more than 1 mm. However, neurites are vulnerable to hypoxic conditions which have to be expected for the 1 mm nerve segment within the hole. It may be speculated that this was the reason for lacking success in biological regeneration experiments. Nowadays, most of the perforated substrates are constructed with the holes and the electrodes to be situated on a very thin membrane (4±10 mm) which is held by a thicker surrounding frame. The frame (usually 100±300 mm thick) ®rst stabilises the membrane, and secondly carries pre-processing circuitry, bonding pads for external connections, etc. In addition, other components of the implant, e.g. nerve guidance channels, are ®xed on the frame. These nerve guidance channels also must permit regeneration of axons, particularly by allowing supply of the ®bres with oxygen and nutrients. Despite the good theoretical understanding of the demands on a device for regenerating nerve ®bres, the big progress in microfabrication and microelectronics and the hard experimental work on this ®eld, the real break-through seems not to have been achieved until now. Reliable nerve regeneration through the holes still is not guaranteed, the researchers often encounter mechanical problems with the implant (broken connections, etc.), and the signals recorded ®nally by the device often turn out to be of non-neuronal origin. The microsurgical technology for the implantation of electrodes applies to peripheral nerves whose axons regenerate and can pass through the o€ered holes of perforated chips. Since regeneration of central nerves is possible only in particular sensory systems, e.g. optic and olfactory nerves of animals under conditions of grafting, the contemporary electronic devices may be ®rst optimised within the peripheral nervous system only. Besides the basic advantage to regenerate, peripheral nerves are wellaccessible to surgery, are round and thus adjustable to electronics, and the success of implantation is easier to investigate. In contrast, most of the central nerve pathways are hidden within the skull or spinal cord and intermingled with neighbouring pathways, thus becoming inaccessible to the microsurgery mentioned above. The only example of a central nerve which may be the target for such a surgery is o€ered by the optic nerve whose extracerebral portion becomes of increasing importance in the regeneration research. According to our current knowledge, the optic nerve may receive key function in the attempt to transfer the microtechnology from peripheral nerves

into the CNS. One possible approach is sketched in Fig. 5. It is based on the optic nerve regeneration model which was established in the eighties (VidalSanz et al., 1987) and was studied extensively during the last years (Thanos, 1992; Thanos and Mey, 1995; Thanos et al., 1993, 1996, 1997). After cut of the optic nerve, the stump of the optic nerve is connected with an autologous peripheral nerve graft which permits the axons of the retinal ganglion cells (RGCs) to regenerate and leads the regenerating axons into their natural area of destination, e.g. the superior colliculus or the thalamus. Vision of the animal could be restored partially, as could be shown by the restoration of the pupilloconstriction re¯ex (Thanos, 1992) and by behavioural and electrophysiological experiments (Thanos et al., 1997). In ongoing studies, a perforated MEA is placed between the optic nerve stump and the peripheral nerve graft, i.e. directly into the path of the regenerating axons. By this way, it should be possible to record nerve signals of a part of the surviving and regenerating RGCs with connections within the superior colliculus. The ®rst results indicate that RGC axons grow through such a perforated implant and encourage to use these implants for long-term biocompatibility and recording experiments.

Fig. 5. Scheme of proposed application of a MEA for regenerating nerves in the CNS. The MEA is placed into the path of the regenerating optic nerve. Axons grow through the holes into the peripheral graft and may reach the superior colliculus. In order to visualise the surviving RGC and their axons, the lipophilic ¯uorescent dye 4-Di10-ASP (Molecular Probes, Eugene, OR) was placed at the site marked by the asterisk. The dye was then transported retrogradely into the retina as indicated by the open arrow. The photograph shows surviving RGC and their dendrites as well as several axons running across the retina. The MEA was made from perforated polyimide with thin platinum electrodes (whole thickness 10 mm) and was kindly provided by J. U. Meyer and T. Stieglitz, Fraunhofer Institute for Biomedical Engineering, St Ingbert, Germany.

Implantable Bioelectronic Interfaces for Lost Nerve Functions

3. BIOCOMPATIBILITY 3.1. General Remarks The major prerequisite for the application of implants, e.g. neuroprostheses, is that the organism accepts the implant, i.e. that the implant is biocompatible. It is widely accepted to de®ne biocompatibility as ``the ability of a material to perform with an appropriate host response in a speci®c application'' (Williams, 1987). This broad de®nition comprises aspects of biological, chemical and physical properties of the implant which will be addressed in this chapter. Nevertheless, the concept of biocompatibility is still disputed and depends strongly on the application ®eld of implants. Whereas in some cases surface (chemical) composition of the implant is an important parameter, in other cases physical properties (size, shape, sti€ness) are the major determinants of biocompatibility, particularly under the in¯uence of locomotion (Boss et al., 1995). An implantation is always a traumatic intervention. However, one important way to minimise the consequences is a high biocompatibility of the implant. An implant can be considered to be biocompatible if . it does not evoke a toxic, allergic or immunologic reaction, . it does not harm or destroy enzymes, cells or tissues, . it does not cause thrombosis or tumours, . it remains for a long term within the organism without encapsulation or rejection. For a long-term stable neuroprosthesis, the whole implant must have mechanical and geometrical properties which ®t to the site of implantation in order to minimise traumatic lesions. The concept of biocompatibility is not limited to non-toxicity, but encloses physical and chemical surface properties and whole behaviour of the implant in its biological environment as well. Therefore, two areas have to be considered, the ``biosafety'' and the ``biofunctionality''. Biosafety means that the implant dos not harm its host in any way, and biofunctionality means that the implant acts in the body as it was intended. In addition, ``biostability'' is important which means that the implant must not be susceptible to attack of biological ¯uids, proteases, macrophages or any substances of the metabolism. For example, implants may be subject to continuous attack by hydrolytic enzymes (Salthouse, 1976) or free radicals produced by monocytes and/or cell lysis. Stability of implanted material is important not only for stable function, but also because degradation products may be harmful for the host organism. Overviews about biological reactions to implanted materials have been given, e.g. by Hench and Ethridge (1982), Anderson (1988), Tang and Eaton (1995), and Ratner et al. (1996). Once implanted, a neuroprosthesis has to remain within the body of the patient for many years. In some cases, it is intended to cover the whole life of the patient. This is in particular obligatory in implants for regenerating nerves, because they would have to be cut again in order to remove the

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implant. This implies extreme demands on stability of function and set-up of the implant. Where necessary, e.g. in muscle tissue or in joints, wires and substrates have to be ¯exible enough to allow multiple bending without damage of surrounding tissue and without breaking at the end. Sharp edges which could damage cells and tissues or rough surfaces which allow attachment and growth of microbes have to be avoided. No corrosion is allowed, and insulating polymers must keep their insulating properties. The implant must remain in its implantation site, thus a reliable ®xation must be guaranteed, e.g. by appropriate suturing. In addition, used polymers may not release any substances, e.g. monomers or oligomers, modi®ers, sterilising agents like ethylene oxide, etc. In addition to the implant itself, also materials used for the ®xation of an implant must be biocompatible, as addressed in the perspectives of brain implants by Mo®d et al. (1997). The exception where release of substances is intended are so-called drug delivery systems. Antibiotics, hormones and growth factors could be released in order to prevent sepsis and to improve wound healing, tissue repair and nerve regeneration. The materials of these systems also must be biocompatible, with particular emphasis on preventing protein adsorption and platelet adhesion which could hinder the substances to be released (Park and Park, 1996).

3.2. Implant Materials It is obvious that requirements towards implants are very high. Nevertheless, there are many materials (metals and polymers) which meet these requirements at least to a big portion. Nowadays, platinum is the electrode material of choice, because it is stable and inert. The amount of platinum ions released into the surrounding tissue may be neglected even after long terms of stimulation. During the last years, iridium has been of increasing importance because a stable oxide ®lm can 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 stimulating electrodes. Carbon ®bres or glassy carbon are also used as electrode materials, and they are biocompatible and stable, though they have a higher roughness than metals. Platinum and iridium are established materials in microelectronics, and also carbon can be deposited onto microelectronic structures. As recently reviewed by RÏõ hova (1996), polymers are used as carrier material and for encapsulation purposes. Most common materials are epoxy resins, polytetra¯uoroethylene (PTFE, Te¯on1), silicone rubbers and polyimide. These polymers are biocompatible, electrically insulating and stable. The bulk properties of the polymers can be modi®ed to a certain degree, and also surface modi®cation procedures are performed in order to improve biocompatibility.

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3.3. Response to Implantation When an implant is brought into the body, the ®rst event is that proteins adsorb onto the surface of the implant. A dozen proteins can be found in biological ¯uids at concentrations higher than 1 mg/ml, and certainly they will form major parts of layers formed on the implant at least in the initial state of adsorption (Andrade and Hlady, 1987). The details of this process depend on the surface of the implant, the composition of the biological environment and the nature of the adsorbed proteins. Adsorption of proteins such as collagen or ®bronectin can favour adhesion of tissue cells (Seeger and Klingman, 1988; Drumheller et al., 1994). An encapsulation of the implant by autologous material (astrocytes, protein layers, endothelial cells, ®broblasts) is desirable in order to integrate the implant into the organism and ``mask'' it in order to avoid undesired reactions of the immune system, thus promoting incorporation and acceptance of the implant. On the other hand, it was reported that ®bronectin and ®brinogen can enhance the adhesion of di€erent bacteria, e.g. Staphylococcus aureus, which is a common cause of infections after implantation (Vaudaux et al., 1984, 1995). Adhesion of the uropathogen Pseudomonas aeruginosa B4 onto polystyrene was increased when urine-derived a-1-microglobulin was pre-adsorbed (Wassal et al., 1995). In¯ammatory reactions can be another consequence of an implantation. In¯ammations involve vascular, neurological, humoral and cellular responses, and the following acute-phase response is characterised by stress-induced changes in the neuroendocrine and immune systems (Kushner, 1982; Khansari et al., 1990). During an acute-phase response, concentration of di€erent serum plasma proteins increases signi®cantly, such as serum amyloid A, C-reactive protein, ®brinogen, a-1-antichymotrypsin, complement protein factor B and C3, haptoglobin and a-1-antitrypsin, whereas concentration of other proteins is decreased (albumin, transferrin and transthyretin) (Baumann and Gauldie, 1990). Furthermore, leukocytes adhere to blood vessel walls and migrate into the tissue which requires highly speci®c interactions mediated by selectins, integrins and inmmunoglobulins (Jones et al., 1996). Cytokines are released, e.g. IFN-g, TNFa, IL-1 and IL-6, which are principal mediators of the acute-phase response (Baumann and Gauldie, 1990). High concentrations of TNF-a and IL-6, e.g. may induce tissue damage, up-regulation of surface adhesion molecules and enhanced production of proteases and free radicals by macrophages. Another issues of an in¯ammation are fever and in®ltration of monocytes/macrophages, eosinophils, neutrophils, lymphocytes, granulocytes, ®broblasts and giant cells. When an implant is brought into the brain, astrocytes can be observed to respond quickly to this injury as they react to every damage (Cavanagh, 1970; Ludwin, 1985). They proliferate in the vicinity of the implant and send their processes towards the implant. Microglial cells are also activated and transform from rami®ed to amoeboid type. The implant is coated with a layer which contains col-

lagen, giant cells, nerve ®bres in di€erent states and blood vessels (Schultz and Willey, 1976). In order to evaluate biocompatibility of a material, in vitro experiments play the major role (Klein et al., 1995; Hanks et al., 1996), although they do not re¯ect the whole complexity of the in vivo situation. In vitro experiments are based on cell lines and on primary cell cultures depending on the intended site of application. After bringing the cells in contact with the material, di€erent parameters are evaluated, as morphological and ultrastructural changes, cell adhesion, release of mediators, presentation of molecules, changes of metabolism, etc. One of these parameters is metabolism of arachidonic acid by macrophages as indicator for in¯ammatory processes (Charissoux et al., 1996). Release of cytokines is also a measured parameter, e.g. of TNF-a (Hunt et al., 1996). Although neuroprostheses are not applied inside blood vessels, blood compatibility is also an important issue, with particular emphasis on behaviour of blood cells, extent of complement system activation, platelet activation and clot formation. The ability of endothelial cells to produce antithrombotic substances has to be examined (Cenni et al., 1993). There are di€erent cell surface molecules which are important in the context of implants. Endothelial cell cultures are preferred for biocompatibility studies of materials for vascular prostheses, and proteins important for adhesion are expressed upon contact between the implant and the cells. The occurrence of such proteins is studied in biocompatibility evaluations with endothelial cells. Examples are the platelet endothelial cell adhesion molecule-1 (PECAM-1, CD31), the endothelial leucocyte adhesion molecule-1 (ELAM-1), the intercellular adhesion molecule-1 (ICAM-1, CD54), and the vascular cell adhesion molecule-1 (VCAM-1) (Cenni et al., 1995). One aspect of biocompatibility for neural prostheses is damage by permanent charge injection. Permanent electrical stimulation can cause damage of neural tissue, such as gliosis, calci®cation of neurons and other cells, lipid inclusions, or glycogen granules in astrocytes, and neural tissue may be lost ®nally (see Agnew and McCreery, 1990a). The basic parameter for stimulation intensity is the relation between charge density and injected charge per phase. Charge density is measured in mC/cm2 and is determined by the frequency, the applied current and the size of the electrode. Absence of neural damage can be expected for the range of a charge density of 100 mC/cm2 with 0.2 mC/phase or a charge density of 15 mC/cm2 with 8 mC/phase (Agnew and McCreery, 1990a, p. 229). Agnew and McCreery (1990b) came to the conclusion that damage of neurons and axons has its origin in their hyperactivity due to severe electrical stimulation, particularly if stimulation is performed with a high frequency. That is why frequency of stimulation should be as low and pulse duration should be as short as possible. Furthermore, stimulation should not be performed continuously but with cut-o€s where possible.

Implantable Bioelectronic Interfaces for Lost Nerve Functions

3.4. Surface Modi®cation In order to enhance biocompatibility of implanted materials, reduce macrophage adhesion onto the implants and prevent in¯ammatory reactions, surface modi®cations of materials intended for implantation are widely studied. These investigations are performed particularly with polymers because they are the main material used for housing, encapsulation and insulation purposes. Whereas attraction and adhesion of macrophages and other white blood cells to an implant can favour in¯ammations, inhibition of such an adhesion would be an important factor of biocompatibility. Nevertheless, it depends on the intended use of the implant whether adhesion of proteins and cells is desirable. The best example are vascular grafts which require a completely di€erent cell behaviour on their surfaces: inside the grafts minimal cellular adhesion and ®brin formation, but extensive adhesion of tissue cells and matrix formation on the outer side. With a neuroprosthesis, also di€erent properties are needed. Recording electrodes should remain bare for good sensitivity, and the best would be an intimate and stable contact between the electrode and neuron cell bodies or axons. Housing and encapsulation materials would be allowed to be coated by tissue cells like ®broblasts in order to incorporate them into the body. Stimulating electrodes are not coated if the applied currents are high enough to cause Faradaic processes at their surface like oxygen evolution, and processes of in¯ammation, gliosis and neuron damage occur not actually on the electrode surface, but nearby the electrodes. 3.4.1. Surface Roughening First kind of surface modi®cation which can be done is the preparation of smooth or rough surfaces. If minimal cell adhesion is intended, surfaces are made as smooth as possible. If cell adhesion is wished in order to achieve good incorporation into the body, rough surfaces are advantageous. Roughening is performed by particular production techniques, where roughness is an intrinsic property of the material, or it can be done by micromachining techniques, where particular patterns are applied to the smooth material, e.g. by photolithographic or micromechanical procedures. In many cases, Vshaped grooves are milled or etched into the material in order to align growth of ®broblasts and to ®x the implant within the tissue. One important example are dental prostheses, where reliable attachment is essential for long-term clinical success. Long grooves running perpendicularly to direction of insertion of dental implants could impede downgrowth of epithelial cells on the implant, which allows connective tissue to adhere and leads to a better ®xation of the implant (Chehroudi et al., 1991). For many years, carbon has been used for di€erent kinds of implants, such as dental implants, percutaneous devices, tendon and tracheal substitutions and 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

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to improve properties of the latter (Pizzoferrato et al., 1993; Cenni et al., 1995). Main modi®cations of carbon used for implantation purposes are pyrolitic carbon and glassy carbon. Both modi®cations have a turbostratic structure which leads to a microscopically heterogeneous surface distribution of electron density and a microscopic roughness (Jenkins et al., 1972; Wege, 1984; Oberlin, 1989). One possible reason for the observed good cell adhesion properties is the combination of hydrophobic surface which favours protein adsorption and roughness which gives cells good ``points of attack'' for adhesion. 3.4.2. Chemical Modi®cation Another kind of surface modi®cation is to change the chemical composition. Ratner (1997) divides these surface modi®cation procedures into biological and non-biological methods depending on the kind of molecules bound to the surface: . non-biological modi®cation: functional groups (amine, hydroxy, carboxy), sulphonates, n-alkyl chains, hydrogels, polymers such as poly(ethylene oxide), poly(2-hydroxyethyl methacrylate), poly(nvinyl pyrrolidone), polyacrylamide, and . biological modi®cation: coating with heparin, hyaluronic acid, sugars, peptides, lipids, enzymes or growth factors. Naturally, layers created by di€erent methods must be stable unless not being intended for biodegradation. After their fabrication, no reactivity may retain, for example by non-saturated binding sites, e.g. sites activated with carbodiimide) or reactive double bonds (e.g. after cross-linking with aldehydes). The well-known statement that ``water is the most biocompatible substance we know'' would lead to the conclusion that hydrophilic and therefore wettable surfaces should be advantageous. Several substances are hydrophilic per se, e.g. glass or silicon dioxide. In other cases, hydrophilic surfaces can be obtained by polar or ionic groups. Such functional groups can be placed on the surface by chemical modi®cation of surfaces or by coating with appropriate molecules, e.g. with proteins or sugars. With a polymer, monomers can be applied for the preparation which carry ionic or polar groups, or copolymers can be prepared in order to combine di€erent properties for desired mechanical properties and surface chemistry (see, for example, Kishida et al., 1991; Arshady, 1993). A surface which has to become hydrophilic can have anionic properties by introduction of negatively charged groups, particularly carboxy (0COOÿ) groups. Neutral hydrophilic surfaces can be obtained by hydroxy (0OH) or amide (0CONH) groups, and cationic hydrophilic surfaces are made by introduction of di€erent amino groups (0NH2, 0NHR0, or NR3). Yun et al. (1995) modi®ed PTFE surfaces with di€erent functional groups and found that cell adhesion and cytokine release were inhibited at best when modi®cation was performed with amide groups to obtain neutral hydrophilic surfaces. Coating of hydrophilic surfaces with polysaccharides, e.g. dextrans, has been recognised to be advan-

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tageous because each monomer within the chain carries up to three hydroxy groups. Indeed, non-speci®c protein adsorption can be reduced (OÈsterberg et al., 1993; Marchant et al., 1994). In order to prepare hydrophilic surfaces and to prevent undesired protein adsorption, surfaces were coated with poly(ethylene oxide), heparin and albumin, and a reduced thrombogenicity accompanied with a reduced plasma protein adsorption and platelet adhesion was found (Amiji and Park, 1993). The authors explain these e€ects by a steric repulsion mechanism because the surface molecules can be regarded as entropic ``springs'' (Milner, 1991). A special technique for the deposition of protein molecules is cross-linking using glutaraldehyde or carbodiimide as shown for arterial prostheses coated with cross-linked gelatine (Drury et al., 1987; Bordenave et 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 crosslinking is derivatisation of molecules to be deposited with photodimerisable groups, such as thymine, cinnamate or coumarin. Upon irradiation with UV light, these groups bind, and derivatised molecules are deposited onto the surface, as demonstrated for the deposition of chondroitin sulphate, hyaluronic acid and gelatine onto Dacron or polyurethane (Kito and Matsuda, 1996). A di€erent kind of introducing hydrophilic groups into a surface is treatment with plasma in a glow-discharge apparatus (Yasuda, 1985; Moroso€, 1990; Piskin, 1992). The plasma can be made of di€erent substances, e.g. from oxygen, ammonia or water, but also from organic molecules such as methacrylates or siloxanes, and a big variety of materials can be treated, even materials which are normally inert, such as polystyrene or PTFE. On such inert materials, plasma treatment is also used in order to form reactive groups for further binding of other molecules. Bigger molecules like polyethylene glycols can also be deposited, and resulting layers resist protein adsorption and cellular attachment, as could be shown for tetraethylene glycol dimethyl ether (LoÂpez et al., 1991). Plasma polymerised siloxanes were shown to provide an anti-thrombogenic surface which is important for vascular grafts and oxygenator devices (Hu et al., 1997). 3.4.3. Speci®c Modi®cations It has been said already that adsorption of proteins onto implant materials is desirable in those cases where tissue cells may adhere in order to facilitate wound healing and incorporation of the implant. Adhering proteins serve as a kind of ``template'' to recruit healthy cells from intact tissue around the site of damage. If the electrodes are implanted into nerves, their appropriately designed surfaces should incite the neurons to adhere to the electrodes and to regenerate their axons along the electrodes. First investigations for the purpose of peripheral nerve regeneration were performed with biodegradable polymers derived, e.g. from extracts of the extracellular matrix (Yannas et al., 1987; Aebischer, 1988; Chang et al., 1989).

For a rational design of regeneration-promoting surfaces, it is necessary to ®nd out the key structures which give rise to the desired behaviour of cells, e.g. neuron adhesion and axonal regeneration. One example of such a key structure is the tripeptide sequence RGD (Arg-Gly-Asp) which is present in ®bronectin, 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 importance of the RGD sequence is also underlined by the fact that it occurs in other adhesive proteins too, e.g. laminin (Grant et al., 1989), collagens, ®brinogen, vitronectin, von Willebrand factor (Hynes, 1987), entactin (Chakravarti et al., 1990), albolabrin, rhodostomin and other viper venom proteins listed in 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 to reduce the big protein molecule (molecular weight of 500,000) to a small tripeptide (molecular weight of 292) with the relevant function. Various materials were modi®ed by di€erent methods with the RGD peptide, as shown for glass, polyethylene terephthalate (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 et al. (1994), for cross-linked polymer networks by Drumheller and Hubbell (1994), and for poly(vinyl alcohol) by Sugawara and Matsuda (1995). Whereas the non-modi®ed materials showed only poor cell adhesion, it could be observed in a high degree after coating with RGD peptides, and also cell spreading and migration were observed in some cases. This indicates that modi®cation of implant surfaces with ``biologically inspired'' synthetic molecules would be able to promote incorporation of the implants into the tissue. Other peptide sequences in proteins had been also identi®ed to be important for cell attachment. In ®bronectin, six additional adhesion sequences have been found besides RGD which are listed in Mooradian et al. (1993). In thrombospondin-1, the sequences RFYVVMWK and IRVVM were found to support attachment of cells (Kosfeld and Frazier, 1993). In the C-reactive protein, the adhesive sequence FTVCL was found (Mullenix et al., 1994). However, there are only few cases of utilisation of these peptides. The ®bronectin-derived sequence WQPPRARI was immobilised on polystyrene and PET, and enhanced adhesion and spreading of endothelial cells was found on the resulting surfaces (Huebsch et al., 1996). Laminin is probably the most investigated ECM protein. It is particularly interesting because it is an abundant component of the basement membranes during development of the embryonic nervous system, but also present in the mature nervous system with apparently important functions not only restricted to guidance or adhesion. In the developing and maturing central nervous system (CNS), laminin plays a crucial role, e.g. in cell migration, di€erentiation and axonal growth (Martin and Timpl, 1987; Kleinman et al., 1987; Martin et al., 1988). Besides the characterisation in vivo, it has been

Implantable Bioelectronic Interfaces for Lost Nerve Functions

extensively used as a substrate for studies of the growth of neurons in vitro. Laminin is a big, multidomain protein (Beck et al., 1990) with many binding sites for di€erent cell receptors (Castronovo, 1993; Mercurio, 1995; Gullberg and Ekblom, 1996; Wei et al., 1997). Amino acid sequences of laminin which have been identi®ed to be important to cell adhesion or growth are, e.g. CDPGYIGSR (Graf et 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), KQNCLSSRASFRGCVRNLRLSR (Gehlsen et al., 1992), YFQRYLI (Tashiro et al., 1994), SIYITRF, IAFQRN and LQVQLSIR (Nomizu et al., 1995). However, only few of these peptides have been used for the modi®cation of surfaces until now. The peptides used in most cases are YIGSR and IKVAV or longer versions of these sequences. Massia and Hubbell immobilised GYIGSR onto glass (1990) and PET and PTFE (1991) and obtained adhesion and spreading of human ®broblasts. Hirano et al. (1993) coupled YIGSR and YIGSR-NH2 to poly(ethylene acrylate) and obtained slightly enhanced attachment of di€erent cell lines. In these cases, e€ects of YIGSR sequence were slightly lower than those of RGD. GYIGSRY was coupled to poly(ethylene glycol) in a cross-linked network, and good ®broblast adhesion was achieved vs. no adhesion on bare polymer (Drumheller and Hubbell, 1994). Fluorinated poly(ethylene propylene) was modi®ed with YIGSR and an IKVAV peptide with 19 amino acids (19-mer IKVAV), and attachment of neuroblastoma and PC12 cells was evaluated (Ranieri et al., 1995). In most of these cases, it was reported that albumin pre-adsorption was necessary for ecacy of immobilised peptides, probably because adsorbed albumin helped the bound peptide to achieve a relatively natural conformation on the hydrophobic surfaces. O€enhaÈusser et al. (1997) coupled the 19-mer IKVAV peptide onto amine-modi®ed glass, silicon wafers or microelectronic surfaces. Embryonic rat hippocampal neurons attached to the modi®ed surface and developed processes. It could be shown by the patch-clamp technique that these neurons were able to produce action potentials. Another laminin peptide, RNIAEIIKDI, was used by Matsuzawa et al. (1996) to modify glass substrates. Again, embryonic rat hippocampal neurons attached to the modi®ed surface and developed a mature morphology with outgrowing axons similar to that of neurons growing on laminin. In both cases, a chemically de®ned medium without serum was used. This may be possible because the surface is hydrophilic due to its prior modi®cation with amino groups. In our laboratory, we investigate possibilities of modi®cation of electrode surfaces in order to achieve a stable contact between neurons and electrodes. For this purpose, we tested di€erent laminin peptides (SRARKQAASIKVAVSADR and SIKVAV, RNIAEIIKDA, YFQRYLI, CDPGYIGSR, PDSGR, GTFALRGDNPQ) which were prepared by Kienle (1997). The immobilisation method

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should be quick and simple, and, furthermore, should modify only the electrodes, but not the substrate without expensive photolithographic techniques. These requirements directly lead to the technique of electrochemical polymerisation. The peptides were modi®ed with a polymerisable group, 3-hydroxyphenylacetic acid. Immobilisation of peptides by this technique could be demonstrated already for antigenic peptides (Heiduschka et al., 1996, 1997). After immobilisation of the peptides listed above onto glassy carbon, enhanced adhesion of embryonic chicken neurons was found, and neurite outgrowth from both adhered neurons and stripes of retina tissue could be observed (Huber et al., 1998). These e€ects di€ered depending on the peptide, and laminin as whole protein molecule showed better ecacy when immobilised, particularly for the outgrowth of axons out of retina tissue. Best results with synthetic peptides were obtained with the 18-mer IKVAV peptide. It further could be shown that electrochemically polymerised peptide layers on recording electrodes which provide an intimate contact between neurons and electrodes would not insulate the electrode. After polymerisation of 3hydroxyphenylacetic acid, the impedance of microelectrodes did not change signi®cantly (Valderrama et al., 1995). For future developments, it would be desirable to ®nd special peptide sequences to which each type of cell responds speci®cally upon contact with the modi®ed surface. In this case, attachment and repulsion, activation or deactivation of each type of cell could be directed in the desired way as sketched in Fig. 6.

4. AUDITORY IMPLANTS The cochlea implant is one of the ®rst implants which have been developed for practical use, and it has been working well since more than ten years in patients su€ering from auditory impairment. It will be therefore discussed as an example of successful development of a neuroprosthetic implant.

Fig. 6. Summary of optimum surface behaviour of an implanted neuroprosthesis: encapsulation surface (left), recording electrodes (left bottom) and stimulating electrodes (right bottom).

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In a healthy ear, the sound vibrations are collected by the outer ear and sent down the ear canal to the eardrum which vibrates the three small bones of the middle ear. Subsequently, the ¯uid in the snail-shaped cochlea vibrates, and these vibrations stimulate the approx. 30,000 tiny hair cells which are located in the organ of Corti inside the cochlea. The electrical signals produced by the hair cells stimulate the bipolar cells of the spiral ganglion, which central ®bres form the auditory nerve which leads the signals into the brain where they are interpreted as sound. The sound is tonotopically represented in the cochlea: high frequencies (up to 20 kHz) are sensed in the basal region, whereas low frequencies (down to 16 Hz) are sensed in the highest part of the cochlea, the apex. This distribution has its origin in a 104:1 gradient of the sti€ness of the basilar membrane and is the prerequisite of the tonotopic organisation of the auditory system. Besides this site-dependent principle of frequency discrimination which is called spectral analysis, there is another mechanism called periodicity analysis. Here the frequency is ``calculated'' in the brain based on the time period of incoming action potentials. In case of a disease or a trauma, the hair cells may be diminished or damaged. The same may happen at increasing age. A widespread reason for acquired hearing impairment or even deafness is meningitis, where the site of damage is almost always the cochlea with loss of the organ of Corti. As a consequence of hair cell loss, the auditory nerve may not be stimulated, and even the loudest sounds may not be heard. That is why hearing aids which only amplify loudness of sound do not help in inner ear diseases. When the auditory nerve itself still is intact, it may be stimulated arti®cially by electrodes implanted in the cochlea, and this is what is done with the so-called cochlea implant which is applied now for approx. 20 years. The success varies due to heterogeneity of diseases and the di€erent degree of destruction. When the auditory nerve itself is destroyed in addition to hair cells, e.g. due to surgical removal of bilateral tumours of the acoustic canal, a cochlea implant does not make sense any more. In such cases, direct stimulation of the brain areas may be tried in order to elicit acoustic sensations. 4.1. Brain Stem Implants Such a stimulation can be performed by the auditory brain stem implant (ABI), which electrically stimulates the auditory pathway at the level of the cochlear nucleus. For this purpose, it could be shown that penetrating electrodes are more e€ective than surface electrodes, and their necessary stimulation threshold was considerably lower (67.5 vs 11.4 mA), their dynamic range was bigger (13.1 vs 24.5 dB), and metabolic activity of some CNS regions was higher as could be shown by higher uptake of [14C]2-deoxyglucose (El Kashlan, 1991). The correct placement of an ABI may be aided by recording of brain stem responses which are recorded during the implantation (Waring, 1995, 1996). The evoked response generally had 2 or 3

waves, and peak latencies of these waves were approx. 0.3, 1.3, and 2.2 msec (Waring, 1995). McCreery et al. (1992) also utilised evoked potentials to optimise the position of stimulation microelectrodes in the cochlear nucleus in experiments with cats. They used 75 mm iridium wires for stimulation, and neurons and neurophil adjacent to the microelectrodes appeared to remain undamaged except for some small gliotic scars. However, a certain depression of neuronal excitability was found which implies the necessity to ®nd the lowest possible excitation threshold by careful monitoring of both the psychophysical threshold for auditory percepts and the electrically evoked auditory brain stem response. Brackmann et al. (1993) described implantation of a single-electrode ABI into the auditory brainstem of patients who were totally deaf due to removal of their bilateral acoustic neuromas. Correct positioning of the electrode in the lateral recess of the fourth ventricle is very important for maximised auditory sensation and minimised activation of other nerves. Shannon et al. (1993) reported that these implanted electrodes were stable for more than 10 years, and that the auditory sensations produced by the implant were similar to the results achieved by single-electrode cochlea implants. Patients could discriminate sound when it was combined with lipreading. By this implant, tinnitus reduction could be achieved in several patients (Soussi and Otto, 1994). Later on, also implants with eight electrodes have been applied (Otto and Staller, 1995). Twelve patients received such an implant, and eleven of them received useful auditory sensations. 4.2. Cochlea Implants Though there are slightly di€erent views about who is a candidate for cochlear implantation, it is widely accepted that recipients of such an implant should be at least 2 years old, have a severe or a profound bilateral hearing loss and receive little or no bene®t from conventional hearing aids. Successful application of a cochlea implant requires proper training, thus the recipients must have a high motivation. An appropriately working device is expected to allow the patient to detect speech and environmental sounds, to use the telephone, to improve lip-reading abilities, to distinguish between di€erent kinds of environmental sounds and, in case of children, to improve speech and language learning. How does the cochlear implant work? As outlined before, it by-passes damaged parts of the ear and directly stimulates nerve ®bres of the auditory nerve. These signals may be interpreted in the brain as a sound after a period of rehabilitation. The whole device consists of two parts, one situated outside the head (microphone, speech processor, power supply, transmitter) and one implanted into the ear (receiver, stimulating electrodes). The sound recorded by the microphone is converted in a series of electrical signals by the speech processor which is then transmitted through the skin to the receiver (Fig. 7). The signals are led to the electrodes, which apply the signal to the auditory nerve ®bres. The right program-

Implantable Bioelectronic Interfaces for Lost Nerve Functions

Fig. 7. (A) Scheme of a cochlea implant with arrows indicating informatiom ¯ow. Sound is recorded by a microphone and converted into a series of electrical signals by the speech processor. The electrical signals are transmitted through the skin to the receiver which converts them into stimulating pulses applied by the stimulator to the neurons of the auditory nerve. 1Ðear canal, 2Ðeardrum membrane, 3Ðthree little bones (hammer, anvil, stirrup), 4Ð cochlea, 5Ðauditory nerve. (B) Detailed view of the cochlea with the inserted stimulator. The electrodes are distributed along the implant thus providing spatially di€erentiated stimulation. 1Ðscala tympani, 2Ðscala vestibuli, 3Ð basilar membrane with the organ of Corti which contains hair cells and sensory nerve endings, 4Ðganglion spirale, 5Ðauditory nerve.

ming of the speech processor starts 4±6 weeks after implantation, and this may take several weeks to months. The whole procedure of programming is tightly connected with adaptation processes in the brain which may require synaptic plasticity in order to rebuild the auditory system. Although there are 30,000±40,000 nerve ®bres in the auditory nerve, stimulation can be performed for technical reasons by only few electrodes. Commercial devices usually have 6±22 electrodes. Based on the tonotopic organisation of the auditory system described above, it would be straightforward to arrange stimulation microelectrodes along the windings of the cochlea and apply distinct electrical stimulation pulses according to the occurrence of distinct frequencies in the environmental sound recorded by the microphone. The speech processor would analyse the recorded sound with respect to distinct frequency ranges and generate an appropriate series of electrical pulses for every single stimulation electrode. Indeed, this approach is performed with a variety of commercial multi-channel systems and there has been good success upon implantation of such devices. The number of stimulating electrodes mentioned above should be increased in order to achieve a higher quality of stimulation and thus a better acoustic sensation by the patient. However, if

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the number of electrodes is increased, their distance gets smaller, and considerable interference and cross-talk between the channels occur (Hartmann and Klinke, 1990). One way to resolve this problem and to minimise interference was development of new signal processing schemes where only one electrode is activated at a given time. This method is called ``sequential stimulation'' in di€erence to simultaneous stimulation. However, spatial resolution of electrical stimulation cannot be enhanced as it would be desirable. As described already for the brain stem implants, also with cochlear implants measurement of evoked potentials is used to evaluate e€ects of electrical stimulation. Whereas most information can be gained by recording of responses of individual cells or ®bres (Merzenich et al., 1973; Glass, 1983; Hartmann et al., 1984; van den Honert and Stypulkowski, 1984), measurement of evoked potentials is much more convenient, can be performed in alert animals and can be used for the investigation of long-term changes. Among the evoked potentials, recording of the middle latency responses seems to be useful (Kileny and Kemink, 1987; Burton et al., 1989), and extended investigations have been performed, e.g. by PopelaÂrÆ et al. (1993, 1995). The auditory brainstem response (ABR) is often used to evaluate the residual hearing of patients as well as the performance of cochlear implants. In most clinical cases, click-evoked ABRs are recorded. Of course, measurement of evoked potentials is not a hearing test per se. In the ®rst implants, only a single electrode has been used, and provided therefore limited success. Cochlear implants gained higher acceptance since MEAs have been applied which allow multi-channel stimulation. The electrodes are arranged on a whorlshaped carrier which is introduced into the cochlea so that the single stimulating sites are placed near to the appropriate sites of the organ of Corti and the surface of the cochlear nucleus. However, it must be taken into account that damage of components of the cochlea can occur by the surgical process of insertion of the long electrodes applied for multielectrode stimulation. The spiral ligament is ®rst site of damage in these cases, and also the basilar membrane could be disrupted which would lead to a lesion of remaining dendrites and subsequent degeneration of the spiral ganglion cells. This means that potentially existing residual hearing may be destroyed. On the contrary, there are reports that inserted implants caused minimal damage and are well tolerated (Burton et al., 1996). In order to overcome the problem of cochlea damage, several ``soft'' surgical strategies have been developed (e.g. use of lubricating liquids for better insertion, Rogowski et al., 1995) which will not be further discussed here. Cohen (1997) critically contemplates the concept of ``soft surgery'' and states that success of the cochlea implant mainly depends on full electrode insertion, the right stimulation strategy and the survival of a sucient 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 be inserted into the cochlea any more. Drilling is only a

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partial solution because the drilled out portion is limited to the ®rst part of the basal turn, and stimulation results are poor. One possibility is not to insert the multi-electrode carrier into the cochlea at all, but to drill small holes into the cochlea at certain sites and insert small single electrodes through these holes (Chouard, 1994). Although this procedure appears to be delicate and time-consuming, it looks more advantageous, particularly because positions of the electrodes may be set very exactly which favours tonotopic excitation. Another way is the development of an implant consisting of two arrays (Lenarz et al., 1997). Whereas the ®rst array is inserted into the drilled out basal turn, the second array is inserted into the second turn which has to be opened for this purpose. The authors reported that an improved performance of the implant could be achieved. In this context, it was argued that the tonotopic theory for cochlear implants is not valid because dendrites of the spiral ganglion cells may retract or degenerate after loss of the hair cells (which is discussed later), and therefore multi-electrode implants would not be useful. It thus may be concluded that selective stimulation of the spiral ganglion cells would not be possible, and consequently only one electrode with an electrical ®eld penetrating the whole cochlea would be enough. However, it is well established that a tonotopic organisation occurs both in the healthy auditory system and in the case of deafness. Naturally, quality of the tonotopy strongly depends on the time of onset and duration of deafness. In any case, the performance of patients with multichannel cochlear implants is anticipated to be maybe better than the performance of patients with single-channel devices (Gantz et al., 1987; Tyler, 1987). Experiments with deafened cats showed variations in parameters of auditory evoked potentials recorded in individual tonotopical cortical places when the auditory nerve was stimulated with di€erent con®gurations of electrodes through a multi-electrode implant (PopelaÂrÆ et al., 1995). Stimulation with monopolar electrodes yielded results with much greater variability among individual cats and induced profound functional alterations in the CNS (Leake et al., 1995). For this reason, the authors stated that application of monopolar devices should be contra-indicated for at least young children. Comparative studies in humans showed that the outcome of an implant is better with a higher number of active electrodes (Amadori et al., 1996). In 1995, more than 12,000 people world-wide used cochlear implants (for an overview, see, for example, NIH Statement, 1995). A majority of them is able to understand up to 80% of high-context sentences, recognises environmental sounds and can listen to music. In spite of good experience with these devices, there are also limitations. Noisy environment remains a problem. Cochlear implants may not provide the dynamics of sound, i.e. the range of amplitude, like normal hearing does. While normal hearing provides loudness di€erences of 4 orders of magnitude, only a factor of 10 may be heard by such an implant. Another problem is that success of implantation and rehabilitation cannot be predicted

very exactly. Several deaf people still cannot use the telephone and depend on lip-reading, and some of them are helped only slightly by these implants. Failure to surgically recover auditory function is particularly observed when individuals were born with deafness. Poorer results were also achieved in children with pre- or peri-lingual onset of deafness compared to children with post-lingual onset of deafness. However, the di€erence between these two groups appears to lessen with time (NIH Statement, 1995). The reason for this failure is most probably absence of the appropriate synaptic connections in the brain which would allow to interpret the incoming nerve signals as sound. For acoustic sensations, especially with patients who are deaf from birth, a lot of connections have to be established in the brain. This task is certainly easier to accomplish in a developing brain than in an adult one. This may explain the observation that implantation at an age of 2 years ultimately results in a better auditory performance than implantation at the age of 3 years or later (NIH Statement, 1995). Special problems of the application of cochlea implants in children are reviewed by Langman et al. (1996). The observation that individuals with shorter auditory deprivation achieve better results than people with a longer deaf period can be seen in the same context as the ability of the brain to process signals incoming from the auditory nerve. It has been known for a longer time, that sensory deprivation leads to plastic changes in the corresponding areas of central nervous system, particularly in the cortex. In case of removal of sensory input, the deprived area of the somatosensory cortex becomes responsive to neighbouring regions (Kaas et al., 1983; Kaas, 1991). Another problem is that often not only hair cells but also neurons of the spiral ganglia are a€ected by the disease or trauma which lowers the prospects of a good auditory performance of a cochlea implant. There are hints that hair cells in the vestibule may be regenerated (Forge et al., 1993; Warchol et al., 1993), but these ®ndings are controversial (Rubel et al., 1995). Moreover, hair cells in the cochlea are much more di€erentiated, and regeneration of lost hair cells in the mammalian cochlea appears not to be possible at the moment. First e€ect after loss of hair cells is retraction and degeneration of the dendrites of the bipolar spiral ganglion cells. It is generally known about the nervous system that neurites may retract when the target cells are lost because neurotransmitters and other factors are not supplied any longer. In the cochlea, BNDF and NT-3 are produced by the developing organ of Corti, and NT-3 is produced by the inner hair cells also in the adult stage (Ylikosi et al., 1993; Schecterson and Bothwell, 1994). Moreover, transcripts for trkB and trkC which are speci®c receptors for BDNF and NT-3, respectively, were found on the auditory neurones (Ylikosi et al., 1993). Recently it was found that NT-3 can elicit a tropic response on outgrowing auditory neuronal processes in vitro (Malgrange et al., 1996). In the guinea pig cochlea, auditory neurones underwent apoptotic cell death after destruction of associated hair cells, and perfusion of BDNF and/or NT-3 onto the scala tym-

Implantable Bioelectronic Interfaces for Lost Nerve Functions

pani almost completely rescued the auditory neurones from this apoptosis (Staecker et al., 1996; Ernfors et al., 1996). In addition, infused neurotrophins can also have tropic e€ects on the dendrites of the auditory neurones (Staecker et al., 1996). One supplementing measure to improve success of cochlea implants would therefore be to continuously deliver NT-3, at least during the ®rst time after implantation. This could be done by coating the implant with a permeable polymer ®lled with the neurotrophin which then would be released over a certain period of time. On the other hand, experiments with young deafened cats showed that ganglion cell survival can be improved with electrical stimulation mediated by such an implant (Leake et al., 1991), and this neuroprotective e€ect was found to be restricted to the area of stimulation (Leake et al., 1992). However, if the site of implantation and the stimulation pattern are not appropriate, spiral ganglion cell rescue is less e€ective, and functional organisation of the central auditory system may be modi®ed as could be shown in young cats (Leake et al, 1995). Under permanent discussion is which kind of contact between the external electronics and the implanted electrode would be most suitable. Most of the systems apply transcutaneous connection, i.e. the skin remains intact, and the stimulation information is passed through the skin by electromagnetic coupling. This requires a transmitter outside the head and a receiving antenna beneath the skin which gives the information to the electrodes. The other possibility is to let the leads pass directly through the skin. Such a percutaneous system allows an easier troubleshooting and a more ¯exible connection between speech processor and the electrodes. Moreover, they are compatible to magnetic resonance imaging (MRI), whereas transcutaneous connectors utilise a magnet and ferrous materials which are incompatible with the high magnetic ®elds produced by an MRI device.

5. VISUAL IMPLANTS Blindness can be caused by a variety of ocular diseases and systemic disorders. Among them, retinal diseases possess a prominent position as they frequently reduce visual acuity and result in non-curable blindness. Age-related macular degeneration (ARMD) and retinis pigmentosa (RP) are wellknown diseases with high incidence. Approximately 1.2 million people are a€ected by RP which is a hereditary disease and more than 5 million individuals su€er world-wide from ARMD. The result of RP is a progressive degeneration of photoreceptor cells within the retina. Whereas macular degeneration leads to a central loss of vision, patients su€ering from RP ®rst lose peripheral vision which is then accompanied by loss of central vision. With the contemporary technologies, electronic devices will not be able to replace the function of the eye completely. The human retina contains approx. 120 million rods and 6 million cones which converge their signals to about 1 million ganglion cells. For an arti®cial electronic device, a resolution

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of several hundreds of pixels (picture elements, points) can be estimated to be the absolute minimum for a useful optical image which would allow rough orientation, i.e. recognition of doors, stairs, cars, etc. The problem is not to record a picture with a high resolution, but to transmit this information in a meaningful way to the appropriate site(s) within the brain. A huge number of single stimulating electrodes is needed, and di€erent levels of pre-processing of the optical information are required depending on the place of implantation. For this purpose, it is inevitable to understand how the optical information is encoded within the retina and how it is transmitted to the brain to create ingenious mental imagination. The possibilities of application of visual prostheses are illustrated in Fig. 8. Two main types are in development, retinal and cortical prostheses. Retinal prostheses can be applied if the retinal ganglion cells (RGCs) and hence the optic nerve are still intact, and they are thought to replace function of lost photoreceptor cells. If the RGCs are lost, the optic nerve degenerates, and cortical prostheses which electrically stimulate the visual cortex may be developed in these cases. In the following, both kinds of prostheses shall be discussed. 5.1. Retinal Prostheses As outlined above, retinal prostheses are intended to mimic the function of lost photoreceptor cells. The photoreceptor cells are the ®rst in the excitation chain which follows the optical stimulus. Their membrane potential EM is decreased (absolute value, i.e. the amplitude, increases) depending upon extent of light irradiation by decreasing membrane conductivity for Na+ ions gNa, whereas EM is increased by increasing gNa in the case of darkness up to ÿ30 mV. The potential is then synaptically transmitted to the next layers of the retina, i.e. the horizontal cells, the bipolar cells, the amacrine cells, and ®nally the ganglion cells. The RGCs are organised into cells with so-called receptive ®elds of ON± OFF characteristics. The sophisticated signal processing like divergence and convergence within the retina makes it possible to enhance contrast, detect motion, to adapt to di€erent light intensities and match the images into a topographic fashion which is then the basis for binocular vision. In the concept of retinal prostheses, an implanted MEA applies light-dependent electrical pulses to the RGCs which are then transmitted as a series of action potentials to the brain to result there in a meaningful image. There are two di€erent ways to implant such a MEA: (i) ``on top'' of the retina (epiretinal implant), i.e. between the ganglion cell layer and the vitreous (Wyatt and Rizzo, 1996; Rizzo et al., 1996; Humayun et al., 1996; Eckmiller, 1997) or (ii) ``beneath'' the retina (subretinal implant), i.e. directly at the place of the lost photoreceptor cells (Chow and Chow, 1997; Zrenner et al., 1997). With an epiretinal implant, the distance between electrodes and RGCs is smaller, and hence RGCs can be stimulated more directly. In addition, also the axons traversing from more peripheral sites than the location of the implant may come under

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Fig. 8. Examples of investigated possibilities to (partial) restoration of vision by the application of microelectrode 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 replacing photoreceptor cells (subretinal implant), and nerve signals of stimulated RGC are forwarded into the brain. 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.

the in¯uence of the stimulating electric ®elds. Thus, excitation of these axons markedly lowers local speci®city of stimulation at the site of implantation. With a subretinal implant, the electrodes could perform directly the function of photoreceptor cells by giving a stimulus upon irradiation. Subsequent information processing would be performed by the other neuron layers of the retina. This would not be possible with an epiretinal implant, and therefore the latter would require a highly sophisticated encoding of visual information into stimulating signals. Numerous experiments have been performed to achieve excitation of RGC cell bodies rather than axons. This task is complicated because the detailed position of stimulating electrodes relative to the cell bodies cannot be controlled. One possibility is to utilise di€erent excitation thresholds of cell bodies and axons upon cathodic or anodic stimulation, and it seems that preferred excitation of cell bodies can be achieved by anodic stimulation. Another approach is application of non-radial electrode con®gurations. When the lines of the electrical ®eld of linear electrodes arranged in parallel run perpendicularly to the axons, a minimal potential di€erence is built up within the axon, and in the other case, with the electrical ®eld longitudinal with the axon, the potential di€erence would be higher thus lower-

ing threshold for excitation. Currently, geometries of MEAs are under investigation with non-symmetrical electrodes in order to improve selectivity of ganglion cell stimulation. With subretinal implants, stimulation is supposed to be more similar to the natural way because incoming light would be applied as an electrical signal directly at the place where it hits the retina, and the same is done by the photoreceptor cells. Consequently, the concept of most subretinal implants consists of an array of photodiodes which directly supply an electrical pulse upon light exposure. Further processing of these signals would then be carried out by the neuronal layers of the retina, and RGCs would ®nally transmit the visual information to the brain. Naturally, neuronal layers beneath the ganglion cell layer are necessary for this concept. However, it must be expected that they degenerate step by step because photoreceptors which supply stimulating signals are lost. In fact, such a degeneration has been observed in all retinal layers, but it is claimed that remaining neurons should be sucient for the application of a visual prosthesis (Santos et al., 1997). Retinal implants are inserted through the anterior part of the eye, and animal experiments are performed mainly with rabbits. The whole surgery is very delicate because the retina is very thin and soft. For epiretinal implants,

Implantable Bioelectronic Interfaces for Lost Nerve Functions

it is crucial to remove the vitreous completely at the place of implant attachment because otherwise ®rm adhesion cannot be achieved. For this purposes, it is advantageous that the vitreous of the rabbit quickly contracts and forms strands upon trauma, and therefore it can be grasped and removed. The epiretinal implant can now be placed onto the retina. For subretinal implants, the retina is incised, e.g. with a sharply edged canule, and then the inner retina may be separated from the pigment epithelium. The subretinal implant can be inserted now between the inner retina and the epithelium. Finally, the detached parts of the inner retina must be attached. As every neural implant, also a retinal implant must meet very high requirements with respect to its biocompatibility and long-term stability. Moreover, an implant within the eye must be especially compatible and stable because in such a delicate organ an in¯ammation or extensive scar formation would have fatal consequences, and surgical interventions cannot be repeated for several times. It is clear that a retinal implant should be soft, tiny and without sharp edges, particularly when implanted beneath the retina. The shape of the implant must ®t the round eyeball which can be achieved by a ¯exible substrate. Nevertheless, it will certainly be of favour to produce the MEA with a vaulted shape. One aspect already mentioned in the biocompatibility section is the charge density of electrical stimulation signals. Humayun et al. (1994) performed bipolar stimulation of di€erent retinae using platinum wires. They found that surface electrical stimulation could be performed in bullfrog eyes (13 mC/cm2), normal rabbit eyes (19 mC/cm2), and rabbit eyes with outer retinal degenerations (112 mC/cm2). These threshold values are within the range denoted in the biocompatibility section, and the authors conclude that electrical stimulation may be a suitable approach for restoration of vision. Until now, implantation experiments with retinal implants have been performed only with animals. The purpose of these experiments was to evaluate biocompatibility and long-term stability of the implants. Currently, several rabbits with subretinal implants measuring 3 mm in diameter and 50± 100 mm thick are being kept for observation, and it seems that they are well tolerated by the ocular tissue (Zrenner et al., 1997). However, there are only few reports about application of active electrodes in such an implant. The means to evaluate function of implanted electrodes is recording of cortical visually evoked potentials (VEPs). They should appear in a similar manner as if the stimulus would be given by an equivalent optical signal. Humayun et al. (1995) reported that electrical cortical responses could be elicited by stimulating electrodes which were placed in the lens of the eye in rabbits after chemically induced photoreceptor degeneration. Rizzo et al. (1996) also could record VEPs upon electrical stimulation of the rabbit retina. Chow and Chow (1997) implanted bipolar strip electrodes into the subretinal space of adult rabbits. The electrodes were driven by external photodiodes. When the photodiodes were ¯ashed in a remote position, cortical responses were obtained similar to the normal light-induced VEPs produced by the pre-implanted eye. Humayun

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et al. (1996) inserted small electrodes into the sclera of blind patients. Spots of light, so-called phosphenes, could be elicited upon application of short biphasic pulses. Patients who previously had been able to see were able to localise the position of the phosphenes accurately according to the retinal area stimulated, which indicates a certain conservation of the visual map. Two subjects were able to recognise movement of electrodes by detecting movement of the phosphenes, and they also could see two phosphenes when two electrodes were active. This is not surprising since phosphenes can also be produced by irritation of horizontal cells, whose localisation is determined within the retina. Phosphenes are rather ``tactile'' signals, but may be well used as orientation guides upon blindness. One problem is delivery of adequate quanta of energy needed for stimulation. Photodiodes used for retina implants are still too weak, that means voltage produced by them upon irradiation with daylight is still insucient. Therefore, it is tried to overcome this problem by installation of a small infrared laser device which is worn by the patient on glasses. The laser beam directed onto the photodiodes is controlled by a computer-assisted video camera. In one possible design, the photodiodes are situated directly on the MEA, and the most straightforward concept is to combine one photodiode with one electrode. In order to avoid damage to retinal neurons by the laser beam, this design is applicable only for epiretinal implants. In another design, the receiving photodiodes are placed at the edge of the retina, and thin wires lead to the stimulating MEA underneath or above the retina. For the epiretinal implant, di€erent concepts for the encoding of the optic information are developed, e.g. by the group of Eckmiller (1997), particularly trying to consider receptive ®elds of the retina. Finally, a subretinal implant may not disrupt supply of oxygen and nutrients to the retina. Therefore, MEAs have been designed with holes in order to allow exchange of substances (Zrenner et al., 1997). 5.2. Cortical Prostheses When the optic nerve is damaged, no information can be conducted from the eye to the brain. Therefore, reconstruction of visual abilities by direct stimulation of the visual cortex has been supposed (Hambrecht, 1995). Cortical implants consist of a MEA placed under the skull directly on the appropriate side of the visual cortex, and the electrodes stimulate cortical neurons to elicit visual sensations. In fact, patients with such implanted electrodes reported to ``see'' phosphenes. There have been numerous investigations about necessary properties of such an electrode array, and, on the other hand, useful information about the organisation of the visual cortex, i.e. the retinocortical map, could be gained by electrical stimulation experiments. Several experiments were performed to correlate the position of the electrodes with the spatial positioning of the phosphenes in the visual ®eld (e.g. Dobelle et al., 1979). In order to enable the patients to recognise more complex structures, e.g. letters, a matrix of phosphenes must be created by simultaneous stimu-

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lation by many electrodes. Such a matrix was simulated by a monitor covered with an opaque perforated mask, and from the experiments was concluded that a MEA consisting of 25  25 electrodes on an area of 1  1 cm2 should produce a phosphene image on a visual acuity of approximately 20/ 30 provided that the MEA is implanted near the foveal representation of the visual cortex (Cha et al., 1992). Schmidt et al. (1996) implanted 38 microelectrodes in the right visual cortex of a 42-year-old woman who had been blind over 22 years. 34 microelectrodes were able to produce phosphenes, and most of the microelectrodes had stimulation thresholds below 25 mA. Brightness and size of the phosphenes could be in¯uenced by stimulation parameters, and separate phosphenes could be detected by the patient when stimulating electrodes had a spacing of 500 mm. Moreover, the authors reported that six phosphenes could be elicited simultaneously, and they all moved simultaneously with eye movements. The task to provide vision by cortical stimulation appears to be much more complicated because current knowledge about vision is far away from a detailed understanding. Although implantation and maintenance of MEAs in the cortex seems to be easier than in the eye, the questions of appropriate image processing and creation of spike trains for stimulation are highly complicated, particularly for moving images and colours.

6. OTHER NEUROPROSTHETIC IMPLANTS The principal idea of replacing nerve function with implants can be drawn back to the work of Sarno€ et al. (1950), who performed respiration by electrical stimulation of the phrenic nerve. However, it took another two decades till development of biomaterials, electrodes and electronics allowed to make completely implantable devices for functional electrical stimulation (FES). FES is applied for peripheral nerves, and there is a broad ®eld of di€erent clinical applications which will be mentioned in this chapter. Respiratory pacing in case of ventilatory insuciency is performed by electrical stimulation of the phrenic nerve in the thorax using a platinum ribbon electrode which is placed behind the nerve (Creasey et al., 1996), or by stimulation of the phrenic nerve with silastic cu€ electrodes or ``half cu€s'' (Glenn and Phelps, 1985). Energy necessary for stimulation is provided by an attached subcutaneously implanted radio frequency (RF) receiver inductively coupled to an external RF transmitter. Phrenic nerve stimulation was also reported to be performed in children (Girsch et al., 1996). It is a widely applied technique despite some occasional problems, e.g. diaphragm fatigue. There are also devices with an internal energy supply via a battery, where only programming is performed with a computer and a radio transmitter (Mayr et al., 1993). Peripheral nerve stimulation is also used for chronic pain reduction (for a review, see StantonHicks and Salamon, 1997). Nashold et al. (1982)

reported on the implantation of electrodes into upper or lower extremities. Pain relief was achieved in a half of cases in the arms and in a third of cases in the legs. Pain caused by trigeminal neuropathy could be reduced by electrodes stimulating the trigeminal ganglion and rootlets (Meyerson and Hakanson, 1986). Waidhauser and Steude (1994) described electrical stimulation of the trigeminal nerve in the case of so-called atypical trigeminal neuralgia which is characterised by long-lasting burning pain sensations without any pain attacks. Peripheral nerves of the hand were stimulated in order to reduce chronic somatic peripheral nerve pain (Strege et al., 1994). Pain management by stimulation of central areas is discussed below. Another important ®eld is assistance of muscle movement of lower or upper extremities, i.e. legs (Cybulski et al., 1984) or arms. The peroneal nerve was stimulated in order to improve walking abilities of hemiplegic patients, particularly to manage the so-called ``drop-foot'' (Teng et al., 1976; McNeal et al., 1977; Waters et al., 1984; Strojnik et al., 1987). The nervus femoralis and the nervus gluteus inferior were also stimulated by implanted electrodes, and paraplegic patients were reported to obtain ability to stand up from the wheel-chair and even ``walk'' several steps with crutches (Kern et al., 1985). A partial restoration of hand function was achieved in tetraplegic patients by stimulation of the median nerve (Kiwerski et al., 1983). Nevertheless, the large ®eld of neurostimulation for restoration of movement of paralysed patients will not be treated in this paper. Finally, treatment of epileptic seizures has to be mentioned. For this purpose, the vagal nerve is stimulated electrically. Uthman et al. (1990) could achieve reduction of the seizure frequency or decreased duration or intensity of seizures. As side e€ects, they reported a tingling sensation in the throat and hoarseness during stimulation. The stimulation was performed with two spiral electrodes wound around the vagal nerve. In a subsequent report, good e€ects of the electrical stimulation were reported, and the authors state that the patients tolerated the implantation and stimulation well without pain, discomfort, or important changes in their daily activities (Wilder et al., 1991). 6.1. Bladder Stimulation In a healthy organism, micturition is achieved by complex interactions of the somatic, sympathetic and parasympathetic innervation of bladder and urethral muscles (Hoyle et al., 1994). The bladder detrusor smooth muscle is controlled mainly by parasympathetic motor neurons located in the sacral spinal cord segment S3 which are excited by descending preganglionic ®bres. During bladder ®lling, the urethral sphincter contracts to maintain continence. At this stage the parasympathetic signalling pathways and the detrusor are inhibited by sympathetic nerves, and the sphincter muscle is contracted by somato-motoric nerves. During micturition, the detrusor muscle contracts followed instantly by sphincter muscle relaxation. A€erent ®bres ascending from the detrusor muscle provide

Implantable Bioelectronic Interfaces for Lost Nerve Functions

information about the stretching state and therefore about the ®lling of the bladder, and a€erents from the urinary tract contribute to a positive feed-back during voiding. In cases of serious neuropathic voiding disorders, e.g. caused by a spinal cord injury, bladder function is disturbed. In many patients with spinal cord injury, re¯ex pathways responsible for continence still work, but micturition cannot be initiated in a normal way. During the ®rst time after injury, no re¯ectory voiding occurs, and the bladder is weak and atonical. Later on, a stage of low bladder volumes and frequent voidings follows. The most common form of micturition malfunction is detrusor-sphincter dyssernergia, i.e. the detrusor and the sphincter are activated simultaneously instead of alternately (Mahoney et al., 1980), which leads to an incomplete voiding of the bladder and a high risk of infection of the urinary tract. Other abnormal disorders can be too weak contraction of the detrusor leading to incomplete emptying of the bladder due to sphincter muscle convulsion or lesion of detrusor muscle as a consequence of in¯ammations, traumatic nerve damage or surgical transfer of the urinary tract. Multiple sclerosis or arteriosclerosis of brain vessels also can lead to incontinence. Patients with spinal cord injury can control their micturition by re¯ectory initiation of detrusor contraction by tapping or pressing their abdomen. In other cases, catheterisation is necessary. Particularly the latter method is an awkward and inconvenient one. At least partial restoration of the bladder function can be performed by FES of di€erent nerves or muscles. Electrical stimulation to restore normal micturition has been investigated since several decades (for reviews, see Schmidt, 1986; Talalla, 1986; Talalla et al., 1987; Tanagho, 1990). It can be applied directly in the spinal cord (Grimes and Nashold, 1974; Jonas and Tanagho, 1975), the sacral nerve roots (Brindley et al., 1986; Tanagho et al., 1989), the pelvic nerves (Holmquist, 1968; Kaekenbeck, 1979), or the detrusor muscle (Boyce et al., 1964; Halverstadt and Parry, 1975; Magasi and Simon, 1986). Among these possibilities, sacral nerve root stimulation has been the most successful in case of intact e€erent innervation of the detrusor muscle. Stimulation of the sacral root can be performed by surface electrodes (Walter et al., 1989) or by implanted intradural (Brindley et al., 1986; van Kerrebroeck et al., 1991) or extradural (Tanagho et al., 1989) electrodes. Most of the e€erent ®bres for the bladder are located in the ventral branch of the sacral nerve root. However, electrical stimulation of the ventral branch with cu€ electrodes also leads to contraction of the urethral sphincter which leads to hindering of bladder evacuation. Moreover, legs also are moved by the stimulation, or penile erection, sweating and piloerection can occur (Nashold, 1974). The reason for this e€ect is the composition of the ventral sacral roots which contain nerve ®bres innervating several muscles of the legs, the pelvic ¯oor, the urethral and anal sphincter, and preganglionic parasympathetic e€erents innervating the detrusor muscle. Small nerve ®bres need a higher stimulus for their excitation than large ®bres, and that is why excitation

453

of small e€erents for the detrusor is always accompanied by the excitation of large ®bres for the sphincter and leg muscles. There are several attempts to overcome this problem, which are reviewed, e.g. by Rijkho€ et al. (1994). Such an attempt is the post-stimulus voiding technique described by Jonas and Tanagho (1975) which takes advantage of di€erent relaxation times of the striated sphincter muscle (0.4 sec) and the smooth detrusor muscle (13 sec), and a commercially available system based on this principle has been developed by Brindley et al. (1982, 1986). Schmidt et al. (1979) achieved reduction of urethral resistance caused by sacral root stimulation by section or blocking of the pudendal nerve, which, however, is a severe and irreversible procedure. Sweeney et al. (1990) tried to avoid this nerve section by reversible blocking of signal transmission through the pudendal nerve by simultaneous stimulation of the pudendal nerve at a more distal site and subsequent mutual annihilation of the two action potentials upon their collision. This procedure is a sophisticated one and requires implantation of additional electrodes. In experiments with dogs, ThuÈro€ et al. (1982) used di€erent excitation thresholds of large ®bres for the sphincter and small ®bres of the detrusor. A high frequency pulse train with an amplitude sucient for excitation of the large sphincter ®bres was applied to fatigue the sphincter muscle, and a stronger stimulus was then given to activate the detrusor. The principle of sphincter fatigue was also applied by Sawan et al. (1996) who developed an implantable computerised electrical stimulation system for bladder evacuation in animal models (dogs) after spinal cord transection. However, several problems remain with all these methods. Unwanted movement of the legs is not avoided. In general, stimulation of the sacral root can be applied only in patients with a complete spinal cord transection or in patients without pelvic pain sensation. Otherwise, electrical stimulation can cause pain because a€erent ®bres of the ventral roots may also be excited (Schalow, 1989). Therefore, the so-called anodal block technique became more and more attractive which is expected to solve the above-mentioned problems (Accornero et al., 1977; Brindley and Craggs, 1980; Fang and Mortimer, 1991). Nerve ®bre membranes near an anode are hyperpolarised, and action potentials cannot pass this zone with a sucient anodic current. Larger ®bres are blocked earlier than smaller ones, and a selective ®bre blockade can be achieved. Tripolar cu€ electrodes were shown to be useful for both excitation and anodal block. For a systematic and successful development far from rules-of-thumb and empirical animal or clinical trials, the behaviour of the system cu€ electrodeaxon has to be evaluated. A ®rst simple model was presented by Altman and Plonsey (1986). A more detailed study of the volume conductor was presented by Ferguson et al. (1987). Rijkho€ et al. (1994) developed a three-dimensional rotationally symmetrical model and calculated geometric and electrical parameters for an asymmetric tripolar cu€ which would allow selective activation of the detrusor without excitation of the sphincter and the a€er-

454

P. Heiduschka and S. Thanos

ent nerves. The authors state that, however, dorsal sacral roots still would be transected in order to abolish autonomic re¯ex contractions of the bladder and to avoid re¯ex incontinence. Nevertheless, electrodes and stimulation protocols have been applied successfully in patients (Rijkho€ et al., 1997a,b). A di€erent technique to achieve selective stimulation of the bladder is the application of microelectrodes within the sacral spinal cord. Carter et al. (1995) demonstrated that sustained elevation of bladder lumenal pressure without simultaneous activation of the urethral sphincter or leg muscles is possible in the cat by electrical stimulation with four 50 mm iridium microelectrodes. The e€ects of stimulation could be varied by changing the positions of the four electrodes and the stimulus parameters. These experiments were continued in order to evaluate histopathologic and physiologic e€ects of implantation and stimulation (Woodford et al., 1996). Electrodes were implanted into the S2 segment of the sacral spinal cord of a cat in order to excite preganglionic parasympathetic neurons, and bladder detrusor muscle contraction could be achieved upon stimulation. Interestingly, this e€ect was not restricted to the stimulation site of the S2 segment. One problem that still has to be solved was movement of implanted electrodes inside the spinal cord, whereas neuronal damage caused by electrical stimulation was less pronounced. 6.2. Stimulation of Spinal Cord and Brain Electrical stimulation in the area of the spinal cord and the brain, i.e. the CNS, is particularly delicate due to the high complexity of the central nervous tissue leading to the enhanced risk of unwanted damage of nerve function in the course of implantation and stimulation mode and to more complicated patterns of responses which may be elicited by stimulation. Only few examples of stimulation in the CNS have to be presented here, and auditory and optical neuroprostheses have been discussed already. Stimulation of the spinal cord is intended for a big variety of applications, mainly in the case of spinal cord injury, where failure of descending ®bres has to be compensated by stimulation of (pre-) ganglionic neurons. One important application is the control of bladder function by stimulation in the sacral area, and di€erent possibilities of bladder stimulation were described in the text above. Another important application is the reduction of pain sensations (Burton, 1975; Kumar et al., 1991; Broggi et al., 1994; Stanton-Hicks and Salamon, 1997), originating, e.g. from diabetic peripheral neuropathy (Tesfaye et al., 1996), arachnoiditis, perineural ®brosis following surgeries, multiple sclerosis, ischemic pain in the extremities from peripheral vascular disease and angina pectoris (Augustinsson et al., 1995), or loss of limbs by accident or amputation. A ®rst preliminary report on pain inhibition was published by Shealy et al. (1967). The treatment is based on the ``gate control'' theory (Melzack and Wall, 1965) which has its origin in the observation that pain in the skin can often be toned down by gently stimulating the skin around the site of hurt

by soft tickling, brushing or massaging. By these tactile stimulants, inhibitory interneurons within the dorsal horn can be activated, thus suppressing transmission of pain. Moreover, these interneurons are in¯uenced by descending ®bres from the brain too, and lateral inhibition can occur via the large diameter a€erents. The advantage is that these ®bres have a low stimulation threshold and can be easily activated due to their large diameter. Upon stimulation, a sensation is produced in the corresponding skin area called paresthesiae which includes tingling and numb feelings. Usually, stimulating electrode arrays are implanted epidurally, and several anode-cathode combinations can be realised at various spinal levels. As complications, wound infection, electrode displacement or fracturing, and ®brosis at the stimulating tip of the electrode can occur. For this reason, high sterility during surgery and prophylactic antibiotics to prevent infections are required. Nevertheless, implantation of a stimulation system is a safe and quick operation in general which can be performed under general or local anaesthesia and is well tolerated by most patients. For an e€ective pain management, the area of hurt must be covered by paresthesiae (Law and Miller, 1982; Barolat et al., 1991). However, Struijk et al. (1993) showed that not only dorsal column ®bres but also dorsal root ®bres may be activated, which was con®rmed by Barolat et al. (1993) who showed that dorsal roots, dorsal root entry zone, dorsal horn, and dorsal columns are involved in stimulation. Another problem is that motor responses and unpleasant sensations can be caused by stimulus amplitudes which are 40±60% above perception threshold (Jobling et al., 1980; Law, 1987; Tulgar et al., 1993), and thus only a small amplitude window is available for stimulation. If it would be possible to activate selectively only dorsal column ®bres with higher amplitudes, a better pain treatment could be achieved. In order to achieve preferential stimulation of dorsal column ®bres, electrode geometries and electrode combinations were varied. Extended theoretical considerations were performed, e.g. by Struijk and Holsheimer (1996), and they could be correlated with clinical observations (Holsheimer, 1997). In order to reduce pain, also brain areas are stimulated. Brain stimulation is usually taken into consideration for patients in whom other forms of pain treatment have failed. Electrical stimulation in the thalamic nuclei VPL and VPM inhibits the activation of spinal dorsal horn neurons by noxious stimuli, and considerable relief of pain can be achieved at the major portion of patients with very few permanent complications (Young, 1990). Unfortunately, not all patients receive e€ective pain relief, and other stimulation targets such as the K-F nucleus in the parabrachial region of the brain stem have been explored in order to provide pain relief to more patients. Ebel et al. (1996) performed stimulation of the motor cortex. However, the bene®cial e€ects decreased in some of the patients after a good initial phase. Multiple e€orts are undertaken in order to restore normal abilities, e.g. free standing or movement of legs, in case of paraplegia. Rushton et al. (1997)

Implantable Bioelectronic Interfaces for Lost Nerve Functions

achieved movement of parts of legs by electrical stimulation in the lumbar root.

7. FINAL CONCLUSIONS Replacement of damaged or lost nerves by arti®cial implants with recording and/or stimulating electrodes has been a goal of many e€orts since several decades. The construction and application of a neuroprosthesis is a common matter of di€erent ®elds of research, such as neurobiology, medicine, computer science, microelectronics, microtechnology, surface science, electrophysiology and electrochemistry. The high complexity of both structure and function of the nervous system is, however, a real obstacle on the way to apply simple neuroprostheses, whereas really functioning, easy-to-handle and long-term stable neuroprostheses are still not available yet. Nevertheless, implants are applied successfully in a few particular cases, which means that a partial restoration of the natural function can be achieved over a longer time period. These special cases are the cochlea implant and implants for bladder stimulation. However, both prostheses are not real neuroprostheses. Other applications, as pain management or stimulation of the limbs, are still not in the state as broad applicability and reliability could be claimed. In the ®eld of replacement of sensory functions besides hearing by cochlear implants, visual prostheses are the target of ongoing e€orts. Since the visual system is the most complicated sensory one, the natural impediments are particularly large. A satisfactory prosthetic replacement of visual function and microelectronic improvements which will allow parallel processing of detectable photic stimuli. Although the loss of other sensory functions seems on the ®rst view to be of less importance, individuals su€ering from either sensory weakness or loss feel it as a devastating event. Whereas photic detection can be performed by a camera and hearing by a microphone, the problem arises with taste and olfaction with no appropriate sensors until now. The molecular and neuronal mechanisms of taste and olfaction are still not completely understood, and mimicking of such processes by arti®cial sensors was realised in part by only a few laboratory setups. Two problems which were not touched in this review are technical realisation of the ready-toimplant device and data processing. In the past, many approaches failed because problems of encapsulation, connection and stability towards leaking and corrosion. Nevertheless, realisation of an appropriate data transmission and processing is still a big challenge, particularly for sensory neuroprostheses where multi-channel stimulation is required. Onchip electronics, multi-plexing and parallel processing are key features to manage the huge amount of data necessary for a useful stimulation. In order to recognise and classify neural signal patterns, di€erent architectures of computer programs called arti®cial neural nets (ANN) have been created. They are also needed to produce arti®cial spike trains for

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stimulation purposes. The goal of current research of computer science is to choose the best kind of ANN for the single task and to make the ANN as reliable and fast as possible. As the next step, e€orts are performed to make neural processors with adaptable neural net circuits.

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