Ординатура / Офтальмология / Английские материалы / Artificial Sight Basic Research, Biomedical Engineering, and Clinical Advances_Humayun, Weiland, Chader_2007

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Figure 11.1. Illustration of the epiretinal prosthesis.

(epiretinally) and fully integrated with photosensitive imaging and microchip stimulating components. It is capable of sensing and processing environmental visual information and transforming that information into a patterned stimulus on a microelectrode array. Electrical connection to the retinal bipolar and ganglion cell layers is accomplished by positioning the array in close proximity (within 20 m) to the retina. The leads and track lines of stimulating electrodes are individually electrically isolated and coated with biocompatible polymer coatings to protect them from intraocular fluids. When energized, the prosthesis can bypass damaged photoreceptor cells and reinitiate threshold membrane potentials in retinal neuronal cells to restore the visual neural pathway.

The rationale for the retinal prosthesis arises from previous experimental analysis of the morphology and electrical response of retinas damaged by degenerative disease. In these eyes, photoreceptor damage is extensive but transsynaptic neuronal degeneration is limited. Initial research with retinal electrical stimulation in patients by Humayun et al. [2] showed that viable neurons could be electrically excited using small extracellular currents delivered by handheld electrodes placed on the retinal surface. Performed in blind human volunteers in the operating room, these short-term tests showed that retinal electrical stimulation resulted in the perception of light in patients who were blind. More recently, a functioning 16-pixel microelectrode array has been chronically implanted on the retina of a patient by Humayun et al. [3]. In the laboratory, external digital imaging and data processing devices were linked with the internal prosthesis and, when the electrodes were current pulsed, the patient perceived light. He could also discern large print letters and everyday objects, such as a

plate, cup, or spoon, on a high-contrast background. These results lend credence to the important contribution an implantable intraocular retinal prosthesis can have on the long-term quality of life of blind people.

Along the road to development of the retinal prosthesis, a number of significant questions about electrode design and electrochemical interactions between the electrode and the retinal neural tissue must be formulated and tested. Current pulsed prosthesis electrodes must depolarize retinal neurons without damaging nerve cells or surrounding retinal tissue. They must also be mechanically reliable for the lifetime of the implant recipient. For patients to effectively view their environment, the pixel number of individually addressable electrodes should approach 1000, placed in a macular area no larger than 5 mm × 5 mm. The small electrode area and large electrode number cause problems of electrode arrangement, charge density, long-term stability, and biocompatibility. Here we report how electrode materials research and biomimetic modeling of the electrochemical interface between electrode and intraocular fluids define the operating limits of the retinal visual prosthesis.

Electrochemical Reactions at the

Electrode–Vitreous Interface

The Interface

A microelectrode array, implanted in the orb of the eye in close proximity to the retina, is embedded in the vitreous humor, the viscous fluid filling the posterior chamber of the eye behind the lens. The vitreous is a gelatinous electrolyte solution, composed mainly of ascorbate, lactate, collagen, glucose, inorganic salts, and hyaluronate in water [4, 5]. The electrode– tissue interface is the point of contact, and plays a critical role in both the efficacy and the safety of neurostimulatory electrode function. When current is delivered to the electrodes of the prosthesis, a redistribution of charge occurs in the vitreous at the electrode surface. The electrical potential (voltage) across the interface depends on the electrical properties of the electrode material, i.e. conductivity, impedance, and efficiency of electrode charge storage and transfer; and the chemical composition of both phases, i.e. presence of oxide, dissolved metal in solution, etc. To be effective, the electrode must deliver sufficient charge for the cell membrane potential of local neurons to exceed the threshold for neuronal depolarization. For retinal prosthesis, electrode arrays in RP patients, this means injecting 0 05–0 6 C of charge per phase of a biphasic stimulating pulse (50–300 A current amplitudes and 1–2 msec phase duration) [6, 7]. The charge density at the electrode–electrolyte interface with a stimulating pulse is the injected charge per phase (amplitude × duration) per unit surface area (expressed as C/cm2) of the electrode. For high charge storage microelectrodes, the total available energy required for threshold excitation of retinal neurons is lower if the energy is clustered on the surface of small diameter

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electrodes. Because of the unique requirements for high pixel number in a limited active area, the charge density burden of retinal stimulatory electrodes is very high. As a result, the potential exists for charge-related tissue trauma in the area adjacent to implanted neural stimulating electrodes. Interfacial charge density with protracted stimulation of the central nervous system has been implicated in stimulation-induced neural injury. The mechanism for this damage is unknown, but is conjectured to be a result of the passage of current through tissue and neuronal hyperactivity [8]. Mortimer et al. [9] chronically stimulated the cerebral cortex of cats with platinum (Pt) electrodes and biphasic current pulses. They observed a relationship between blood–brain barrier damage in the area surrounding the electrodes and the charge density at the electrode surface. McCreery et al. [10] found that the threshold for damage to the parietal cortex of cats during extended stimulation with activated iridium (IrOx) microelectrodes was the result of synergistic interaction of charge density and charge per phase beneath the implanted electrode. In later work, McCreery and his colleagues [11] were able to link stimulus-induced temporary depression of electrical excitability of the cat cochlear nucleus with prolonged neuronal activity. Depression was not correlated with histological changes near the electrodes, suggesting that the stimulation protocol was above the threshold for reversible neuronal over-activity but below the threshold for irreversible neuronal damage. Neural damage in the cortex of the cat was observed by Yuen et al. [12] after continuous stimulation at a charge density of 100 C/cm2 for 6 hours. Less damage resulted from the application of 40 C/cm2 for 20 hours. Rose and Robblee [13] found that the safe charge injection limit for cathodic to anodic biphasic 0.2 msec pulsing of Pt electrodes in saline was 100–150 C/cm2. Based on Pt dissolution data in saline, Brummer et al. [14] suggested limiting charge densities to < 300 C/cm2 for balanced biphasic pulsing. For iridium oxide microelectrodes in saline, the limits can be extended to 1 mC/cm2 because of the greater charge carrying capacity of IrOx. [15]. Understanding the electrophysiological interactions between prosthesis electrical components and excitable tissue and the electrochemical reactions at the electrode–vitreous interface are the key elements to developing a safe and practical retinal prosthesis insert.

Interfaction Electrochemical Processes

Electrodes charged at a tissue surface interact with ionic species in the fluid medium bathing the tissue. The voltage, time duration, and conductive properties of stimulating electrodes determine the electrochemical processes that take place at the interface. Charged species in the vitreous can realign spatially around the electrode in an attempt to satisfy electroneutrality at the electrode surface (non-faradaic or capacitative charging), or they can bind to the electrode surface and transfer electrons into the vitreous in the form of electrochemical reactions (Faradaic charging). Faradaic charging is reversible if reaction products remain bound to the electrode surface. It is irreversible if products leave the electrode surface and redistribute in the vitreous. Neurostimulatory electrodes can be

divided into two categories: capacitative and faradaic. Capacitative electrodes have a metal core coated with a non-conductive metal oxide surface layer. The oxide insulates the conductive metal, creating a dielectric at the electrolyte interface. Electrons moving through the electrode cannot physically traverse the interface, but can attract or repel ions in close proximity target tissue and initiate an action potential [16]. In contrast, metal electrodes transfer electrons directly across the interface, resulting in faradaic oxidation–reduction reactions in the vitreous and the formation of new chemical species. Guyton and Hambrecht [16] compared the pHs of unbuffered saline solutions after injecting 1 mA monophasic current pulses with 1 mm tantalum oxide Ta2O5 capacitative and Pt electrodes. Pulsing the Pt electrode increased the pH at the cathode and decreased it at the anode. No apparent pH change occurred with similar pulsing of the Ta2O5 electrode. It would appear that capacitative electrodes can offer some advantages over metal faradaic electrodes, i.e. they are more resistant to degradation in physiological solutions and are inert to electrical and chemical interactions with extracellular fluids at the interface. However, for the retinal prosthesis, the size constraints on high numbers of densely packed microelectrodes favor the charge transfer properties of faradaic electrodes. In general, capacitative microelectrodes do not have the charge storage ability of noble metal microelectrodes [17], and their charge storage capacity is not high enough to deliver the charge per phase necessary to activate retinal neurons. Attempts to increase charge storage, such as increasing surface roughness, are often accomplished at the expense of electrode strength and durability. Moreover, techniques for making thin oxide films with high dielectric constants that do not leak current are complicated, especially when applied to micron diameter wire. McCreery et al. [8] compared neural damage in the cortex of cats following 7 hours continuous biphasic pulse stimulation 80–100 C/cm2 with Pt faradaic and Ta2O5 capacitative electrodes. Both electrodes evoked action potentials in the cortex, and both damaged the cortex equally. The capacitor electrodes did not depolarize neural ganglia more effectively or safely than faradaic electrodes.

Metal microelectrodes, driven at interfacial charge densities above safe potential limits, initiate irreversible electrochemical reactions in the vitreous. A number of byproducts of these reactions can be toxic to the tissue surrounding the electrode. Rosenberg et al. [18] reported inhibition of cell division in bacterial cultures grown in media containing electrolysis products from Pt electrode charging in saline. Toxic byproducts are usually the result of anodic oxidation of vitreous solutes. Free chlorine and hypochlorite (both strong oxidizing agents) are formed when saline-based medium is oxidized at the Pt–s saline interface [19]. Electrode charging in body fluids also raises the probability for in situ water electrolysis. H2 and O2 gas evolved during electrolysis [20] can result in local pH changes in the region surrounding the electrode and can be physically damaging to delicate ocular tissue. The electrode itself may be oxidized, resulting in slow dissolution of the metal structure and loss of electrode integrity. Pt oxidizes in saline according to the equation: Pt + 4Cl → Pt Cl42− + 2e−[20]. Dissolution is related to the aggregate positive or negative phase charge injected and can occur

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in charged solutions at potentials below the threshold for electrolysis [21]. For safe, long-term prosthesis operation, stimulatory pulses must be of sufficient amplitude and duration to activate neural ganglia without irreversibly generating damaging chemical byproducts in tissue fluid. Microelectrodes for chronic stimulation must have high charge storage capacities and comparatively high thresholds for initiation of interfacial redox reactions.

Pulse Configuration

In considering the thresholds for harmful faradaic reactions at the interface, the nature of the stimulating pulse is a critical factor. DC monopolar (monophasic) electrode charging can supply the minimum pulse amplitude and duration required for neuronal excitation. However, lengthy DC charging can initiate irreversible reactions in the vitreous via faradaic processes [22], resulting in tissue damage. With constant, prolonged monophasic stimulation, the change in vitreous chemical composition with time is unabated; reaction products accumulate and injury results. In 1961, Lilly [23] proposed an electric waveform that did not damage cortical tissue with long-term stimulation. The waveform passed short duration currents first in one direction and then equally in the other (biphasic). The amplitudes of the biphasic currents were equal, resulting in no net charge injection (no direct current component). A fixed duration intraphase delay, inserted between the polar phases of the biphasic pulse, allowed more time for the transfer of charge to the cortex before charge reversal [24]. With this pulse, ionic changes induced in the interstitial fluid during one polar phase were quantitatively reversed during the reverse phase. Faradaic electrochemical reactions produced products that could not discharge from the electrode before the polarity was reversed. The intended result is the injection of no new species in the solution if the stimulating pulse is biphasic, symmetrical, and balanced.

Electrode Test Apparatus

Prospective prosthesis electrodes should be assessed using experimental protocols that test their electrical performance and stability when current pulses are applied in an electrolyte solution. In this way, pulse parameters can be manipulated within the charge density limits for retinal prostheses and toxic products of interfacial reactions can be measured under electrode stimulating conditions expected to be encountered in real-world retinal implants. An electrolysis cell, a model of the eye, approximating the physiological parameters of temperature, photo excitation, and local electrolyte environment found in the human eye, has been designed and integrated with advanced sensing and bioelectrical stimulating instrumentation for the analysis of the performance of potential prosthesis electrodes. A schematic illustration of the apparatus is shown in Figure 11.2. The eye cell is a jacketed, glass chamber filled with 6–8 ml hyaluronic acid200 g/ml in Ames’ medium. Ames’ medium is a synthetic analogue for

Figure 11.2. Electrode test apparatus.

vitreous humor, especially formulated to preserve retinal tissue and photoelectrical response in vitro [25]. During tests, the cell temperature is maintained at 37 C. Test electrodes immersed in the vitreous solution through a hermetic O-ring assembly are connected to a Multichannel Systems Stimulus generator for application of a pulse stimulus. A Pt plate electrode in the sidearm of the cell, representing the counter electrode placed behind the ear of the transplant recipient, completes the electrical circuit to the generator. The cell is sparged with breathing air (100 ml/min) through an isolated gas flow system, supplying circulating oxygen to the vitreous and allowing for downstream analysis of gaseous products of electrochemical reactions at the electrode–vitreous interface. An inline, Figaro tin oxide combustible gas sensor, positioned after the cell in the gas flow line, detects hydrogen gas evolved at the cathode above the threshold of water electrolysis. This system can be used for acute or chronic steady-state repetitive testing of the kinetics of interfacial electron transfer reactions at the neural electrode–vitreous interface. For these experiments, the limits of charge injection were between 0.1 and 1 mC/cm2.

Pt electrodes have been tested for durability and thresholds of electrolysis using the biomimetic eye cell of Figure 11.1. Pt wire 76 m , fused in insulating glass capillary tubing and polished flush with the glass across the fused end, was immersed in synthetic vitreous and driven with biphasic or monophasic current pulses. Charge densities were kept below 1 mC/cm2. Repetitive monophasic pulses were applied from the stimulus generator. Current amplitudes were 50–500 A for 0.1 msec and 1–50 A for 1 msec pulses. The time between pulses was 16 msec. Steady-state hydrogen production was measured as a function of charge density. The results for hydrogen production for 0.1 and 1 msec pulse durations are shown in Figure 11.3. Thresholds for hydrogen generation correlated with charge density at both pulse durations. The charge density for

itance [26]. They must be mechanically and biologically compatible with their implant location and transfer enough charge to depolarize excitable tissue without triggering significant interfacial faradaic oxidation–reduction reactions. Future demands for higher resolution, high pixel, and small diameter electrode arrays will require the consideration of a variety of materials for electrode construction, including Pt and other precious metals, metal composites, and metal alloys. Micron diameter Pt, Ir, Au, and PtIr electrodes can perform well as high density, low impedance cortical stimulatory electrodes. Pt and its more durable alloy PtIr have a long track record as neural stimulatory electrodes, chiefly because of their biocompatibility, relative resistance to corrosion, and high charge carrying capacities [22]. Iridium can be activated by biphasic cycling of the electrochemical potential in electrolyte solutions to form iridium oxide IrOx on the surface of precious metal substrates [27]. Charge is transferred in IrOx electrodes by stimulus-generated valence changes in the surface oxide layer [15]. IrOx electrodes can deliver a net safe charge of 1 0 mC/cm2 [28] with 0.2 ms biphasic pulses in bicarbonatebuffered saline, and the charge storage capacity is much greater than that of Pt or PtIr [29]. Moreover, it is hard and durable, resistant to fowling and dissolution in physiological solutions. In vivo comparisons of the performance of IrOx and PtIr electrodes [30] demonstrated that they could both safely evoke action potentials with extended continuous microstimulation in cat cortex without notable histological damage. Both electrodes performed well for extended periods at charge densities of 800 C/cm2, but poorly at charge densities above 3200 C/cm2. Weiland et al. [31] examined the histology and the electrical properties of the guinea pig cortex implanted with thin-film iridium oxide electrodes. The cortex was energized for two hours with 5–100 A, 100 msec biphasic pulses. Electrochemical impedance spectroscopy and cyclic voltammetry measurements showed that the cortex and electrodes were both affected by a temporary impedance shift but, histologically, the tissue was unaffected.

Titanium nitride (TiN) is a material for electrode construction that offers good potential for extending the safe stimulation limits beyond iridium oxide or Pt [32]. The electrochemical performance of titanium nitride deposited on silicon bioelectrodes was compared with similarly constructed iridium oxide electrodes by Weiland et al. [33]. Titanium nitride had similar charge injection capability to Pt0 87 mC/cm2 , significantly below the capacity of iridium oxide 4 mC/cm2 .

Operational Stability of Pt Retinal Prosthesis Electrodes

The development of a retinal prosthesis for artificial sight includes a study of the factors affecting the structural and functional stability of chronically implanted Pt microelectrode arrays. The maintenance of electrical integrity between the electrode and soft tissue depends, in part, on initial electrode placement in close proximity to the retina and insulation of prosthesis components from intraocular fluids. Only the electrode surface is in direct contact with the retina and the corrosive salt solution bathing the ocular cavity. The 16-pixel microelectrode retinal prosthesis now implanted in the eye of a retinitis pigmentosa patient [3]