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

Figure 13.4. Electrochemical impedance spectroscopy (EIS) measurement indicates that molded Pt electrodes have lower electrode impedances than those of sputtered thin films.

a frequency scan from 1 Hz to 100 kHz. The EIS results indicate that molded Pt electrodes have lower electrode impedance than those of sputtered Pt thin films as shown in Figure 13.4. Lower electrode impedance is desirable for neural stimulation as it will reduce voltage excursion and lower the power consumption during stimulation.

There is a tendency for charged species to be attracted to or repelled from the metal–solution interface. This gives rise to a separation of charge, and the layer of solution with different composition from the bulk solution is known as the electrochemical double layer. As a result of the variation of the charge separation with the applied potential, the electrochemical double layer has an apparent capacitance (known as the double layer capacitance) [11]. It can be seen that in order to achieve large charge storage capacity, for a given material and cell setup, the real surface area of the electrode is a key factor: the higher the surface area, the higher the capacitance. For high-density microelectrode arrays used in neural stimulation, the geometric surface area is limited by the application, and an effective way to increase the electrochemical surface area is to increase the roughness of the electrode without changing the array size.

The electrode capacitance results along with electrode surface areas are listed in Table 13.1. The electrode capacitance is determined by EIS measurements at

Table 13.1. Electrode Capacitance determined from EIS measurement at 1 kHz of molded Pt Samples.

1 kHz at a very small ac voltage of 10 mV. The results indicate that the molded electrodes have better surface roughness than what are found in thin-film metals. For sputtered thin-film Pt surface on polymer substrate, a capacitance is typically in the range of 10–25 uF//cm2.

Cyclic voltammetry (CV) has been employed to determine the operational potential window (the water window) limited by the H2 and O2 evolution potentials due to electrolysis of water on the cyclic voltammogram. The area estimated by integrating the cyclic voltommogram within the water window indicates the charge delivery capacity of an electrode. Figure 13.5 shows the comparison of cyclic voltammograms of sputtered and molded Pt electrodes. The CV measurements were carried out in PBS solution with a potential scan rate of 50 mV/s. The Pt electrodes presented a well-defined voltammogram with enlarged reduction and oxidation peaks within the water window. Such increased cyclic voltammograms for molded Pt over sputtered Pt thin film suggest increased charge delivery capacities for neural stimulation for both materials. Another advantage of these electrodes is that no visible active peaks were detected in cyclic voltammetry measurements. This implies that no measurable active metals, which may cause corrosion or tissue damage, are present in Pt samples. This is a very important consideration point for all neural stimulation applications.

Most neuro-stimulation applications use a biphasic, cathodic first current pulse. Under such current pulse, the electrode’s voltage response (voltage excursion) is a direct indicator of charge injection capability. For a given electrode at a given pulse current, the lower the electrode voltage, the higher the charge

E (mV) vs Ag/AgCl

Figure 13.5. In cyclic voltammetric (CV) test, molded Pt electrode presented a welldefined voltammogram with enlarged reduction and oxidation peaks and had a higher charge storage capacity. The CV measurements were carried out in PBS solution with a potential scan rate of 50 mV/s.

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injection capability. The voltage excursions were measured under pulse stimulation currents to evaluate the molded electrode charge injection capacity. A programmable multichannel stimulator made by Multichannel Systems was used to generate stimulus. A TDS 3014B oscilloscope along with the WaveStar program was used to measure and record the voltage excursion. The electrodes were stimulated in PBS under a charge-balanced, cathodic first, biphasic pulse current at 100 C/cm2 for Pt at 50 Hz. The voltage excursion curves are shown in Figure 13.6. Under the same stimulation conditions and same geometric surface area, molded Pt electrodes presented much lower voltage excursion for the charge density levels tested than what are found in their thin films. A low voltage with linear voltage excursion is favorable for neural stimulation as it will have less effect from irreversible electrochemical reactions which may produce harmful byproducts and/or cause electrode corrosion.

Electrical stimulation of biological tissue with metal electrodes requires the flow of ionic charge in the biological tissue. This flow of charge can be induced by two mechanisms: capacitive and Faradaic. The Faradaic mechanism of charge injection involves electrode transfer across the electrode–tissue interface. This may induce harmful electrochemical reactions and can cause tissue or nerve damage [12]. Hydrogen/oxygen evolution due to water hydrolysis induced by stimulus will alter pH, causing metal corrosion and possible tissue damage in the electrode/tissue interface [13]. In the case of biphasic pulse, no hydrogen

Figure 13.6. The electrodes were stimulated in PBS under a charge balanced, cathodic first biphasic pulse at 100 C/cm2 for Pt at 50 Hz. Under the same stimulation conditions, nanomolded Pt electrodes (b) presented lower voltage excursion than that of their sputtered thin-films (a).

production is observed up to the maximum deliverable current of 800 mA which is equivalent to a charge density of 200 C/cm2.

Based on the results from tested samples, the molded Pt has improved qualities in all aspects of electrochemical performance which suggests they could be the preferred candidates in neural stimulation applications versus widely used thin film electrodes. The experiments also suggest that the molded electrode surface with preserved nanoparticle morphology has increased the nanoscale surface roughness and gained more surface area than that of a smooth surface.

Conclusions

The described nanopowder molding fabrication technology could create high- aspect-ratio solid metal electrodes and wires. This rather simple process provides a possible alternative fabrication method other than existing technologies. Use of this molding technology to make 3D electrode surfaces (protruding bump-like structures) will increase the electrochemical surface area and make it possible to position the stimulating electrode closer to the neurons. The rough surface due to the nanoparticles incorporated on the electrode surface and increased surface area. Initial experiment results suggest that molded structure increases the charge injection capacity therefore minimizing stimulus threshold. For microelectrodes with high DC resistance, the potential exists for unacceptable heat generation. Lower electrical resistance is also desirable for reducing the heat generation.

Molded Pt electrodes outperform their sputtered thin film counterparts in all aspects of electrochemical properties with lower electrode impedance, higher charge storage capacity, and lower voltage excursion. Batch process is suitable for low-cost mass production of high-density 2D/3D conductive structure. In the foreseeable future, there will be a constant need for highquality neural stimulation electrodes and trace wires, especially when more and more prosthestic devices are in development. The maturation of this electrode fabrication technology could bring broad benefits and impact to all prospective applications.

Acknowledgments. This research was supported by the DOE Lab 01–04 and the Oak Ridge National Laboratory, managed by UT-Battelle, LLC, for the US Department of Energy under contract DE-AC05-00OR22725. Support from the DOE Office of Biological and Environmental Research is gratefully acknowledged. Authors also thank Anat Burger, Christen Smith, Karolyn Hansen, Charlene Sanders, Elias Greenbaum, and Chase Byers for their generous help and assistance during the completion of this work.

Reference

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In these diseases, the “imaging” photoreceptor layer of the retina degenerates, yet the “processing circuitry” and “wiring” subsequent to photoreceptors are at least to some degree preserved. The RP occurs in about 1 out of 4000 live births, corresponding to 1.5 million people worldwide. This disease is the leading cause of inherited blindness. The AMD is the major cause of vision loss in people over 65 in the Western world. Each year 700,000 people are diagnosed with AMD, and 10% of these people become legally blind. Currently, there is no effective treatment for most patients with AMD and RP. However, if one could bypass the photoreceptors and directly stimulate the inner retina with information relaying the visual scene, one might be able to restore some degree of sight.

One important factor affecting this strategy is that the absence of normal signaling from photoreceptors can lead to some progressive degeneration and miswiring of retinal circuitry [1, 2]. This type of degeneration is a general property of neural circuits. Thus, for an electronic implant to properly transmit visual signals to the inner retina, any degeneration of circuitry must not drastically change how these signals are interpreted by the higher brain. This is true in the case of cochlear implants, which bypass degenerated primary auditory sensory neurons; both the nerve and the downstream neural circuitry retain the ability to transmit interpretable auditory information.

Indeed, some first steps have been taken toward the development of an electronic retinal implant. It has been demonstrated that degenerated retina can respond to patterned electrical stimulation in a manner consistent with vision [3–6]. Human patients implanted with an array of 16 4 × 4 electrodes of 0.4 mm in size can recognize reproducible visual percepts with patterned stimulation of the retina [3–6]. The patterns perceived by the patients did not always geometrically match the stimulation pattern, which is not surprising knowing the complexity of the retinal spatial organization. However, the one-to-one correspondence between the perceived and the stimulation patterns gives hope that with some learning and image processing the patients might be able to perceive useful visual information from this type of stimulation [7].

A large percentage of patients with AMD preserve visual acuity in the range of 20/400 and retain good peripheral vision. Implantation would be worth its risk for such patients only if it provided substantial improvement in visual acuity. In contrast, patients with advanced RP would benefit little unless the enlargement of the central visual field was enough to allow reasonable ambulation. Normal visual acuity (20/20) corresponds to an angular separation of lines by 1 min [8], which corresponds to spatial separation on the retina of about 10 m, or in other words, spatial frequency F = 100 lines/mm on the retina. To provide such spatial frequency the stimulus pixels should have a linear pixel density at least twice higher: P ≥ 2F, i.e. two pixels per line (Nyquist sampling theorem). In other words, to resolve two white lines at least one black line should be located in between. Thus the maximal spacing between pixels that will allow for resolving two lines separated by 10 m is 5 m. Similarly, spatial resolution corresponding to visual acuity of 20/400 corresponds to a pixel spacing of

about 100 m, while acuity of 20/80 (enough for reading with some visual aids) requires pixels smaller than 20 m. For these estimates, it is understood that retinal stimulation by one electronic pixel may not produce a perceptual pixel-like “phosphene,” and may generate more complex perceptions dependent on the precise number and connections of stimulated cells. What is essential in this analysis is the fact that pixel density determines maximal amount of information or maximal spatial resolution that can be provided by the stimulating array, and thus the best possible visual acuity, if the brain will be able to utilize all this information. Encoding the information, i.e. conversion of the image from the video camera into the map of stimulating signals, is a separate issue.

It has been previously estimated that 625 pixels can suffice for minimally resolving images in a tiny (1 7 or less) central field [9]. For functional restoration of sight a retinal implant should ideally cover a larger field of view – at least 10 (3 mm in diameter), and support a visual acuity of at least 20/80 (corresponding to a pixel size of 20 m and density of 2500 pix/mm2) in the central 2–3 of stimulating area. We must emphasize that these numbers are just rough estimates since functional acuity might be increased due to the eye scanning (hyperacuity) or require less pixels for some tasks through pattern recognition. On the other hand, retinal disorganization and the need for learning the new type of stimuli could necessitate more pixels.

Electrical stimulation of neural cells in the retina has been achieved with an array of electrodes positioned on either the inner[3, 9, 10] or the outer side of the retina [11–13]. Setting the electrodes into the subretinal space so as to stimulate bipolar cells, although surgically more complicated, has the potential advantage that signal processing in the retina is partially preserved. Full utilization of this advantage will probably require intervention at relatively early stages of retinal degeneration, before significant remodeling of the retinal neural network takes place[2]. Exciting the ganglion cells with electrodes positioned on the epiretinal side abandons the visual processing by the inner retinal network directly stimulating the output of the retinal circuitry.

One concern with either technique, pertaining to the goal of high-resolution stimulation, is that the electrodes will always be some distance from the target cells. This occurs because the inner limiting membrane and nerve fiber layer intervene in the case of epiretinal placement, or because of photoreceptor remnants in the case of subretinal implantation. In addition, a diseased retina may have an uneven thickness or wavy structure. Large distances between the cells and closely spaced electrodes result in cross-talk between neighboring electrodes, and the need for a higher charge density and power for cell stimulation. This, in turn, can lead to erosion of electrodes and excessive heating of the tissue. Furthermore, any variability in the distance between electrodes and cells in different parts of the implant will result in variations of the stimulation threshold, making it necessary to adjust the signal intensity in each pixel.

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As shown below [14] for chronic stimulation with pixel density of 400 pix/mm2, which geometrically corresponds to visual acuity of 20/200, the electrodes need to be within 15–20 m of the target neurons. For visual acuity of 20/80, the separation between electrodes and target cells should not exceed 7 m [14]. Thus, ensuring a close proximity of cells to the electrodes is one of the most important unresolved issues in the design of a high-resolution retinal prosthesis. In this article we describe several techniques that may assure proximity of electrodes to the target cells. One of these techniques prompts migration of retinal cells into proximity of stimulating electrodes positioned in the sub-retinal space [15]. During migration the cells preserve axonal connections to the rest of the retina thus maintaining the signal transduction path. Another technique is based on an array of electrodes protruding from the sub-retinal chip [14–16].

A very significant problem with many designs of visual prosthetic systems is that they include head-mounted cameras linked (wirelessly) to the pixels on the patient’s retina, so that eye movements are dissociated from vision. This dissociation greatly compromises the process of natural viewing. When the eye scans a scene, each movement is coupled to a strong expectation that the image will change accordingly. In addition, small eye movements during fixation are actually required for image perception: if an image is stabilized on the retina, it fades from perception within 100 ms [17]. In this article we describe the design of a system with a microcomputer-assisted interface and direct optical projection of the processed image onto photosensitive pixels in the retinal implant using nearinfrared light. This system should allow for natural eye scanning and enable the simultaneous use of implant-stimulated vision and any remaining normal vision at any level of luminance.

Another important aspect of a macular chip design is adjustable image processing. Synaptic connections from foveal photoreceptors radiate out to bipolar and ganglion cells at some distance from the visual center. Thus, an image centered on the foveola will be processed by bipolar and ganglion cells in a circular zone outside foveola. Prosthetic chips will need to have stimulus signals that match this neural anatomy. The system described below includes location-dependent image processing based on a precise tracking system that monitors the location of the implant in real time. Stimulation of neurons by the retinal implant differs from natural retinal signal processing. Therefore, to enable the translation of stimulus patterns into the conscious recognition of objects, visual chips may require some form of image processing and neural “learning”, much as is required by modern cochlear implants. Tracking the implant in real time allows for the position-dependent image processing that may be required to translate visual information into electrical signals that can be properly interpreted by the higher brain.

In the article below, we describe a system that addresses all three issues raised above: (1) proximity of electrodes to the target cells, (2) delivery of information to the retinal implant linked to the natural eye movements, and (3) locationdependent image processing.

Proximity between Electrodes and Cells as a Resolution-limiting Factor

Delivery of high spatial frequencies via thousands of small electrodes is possible only if each of them will be affecting cells only in a close proximity – a zone similar to the size of the electrode. However, simultaneous injection of current from a dense array of small monopolar electrodes having a remote return will result in interference of the local electric fields producing a very far-extending field equivalent to an effect of one electrode similar in size to the whole array, i.e. the field will extend to a distance of about 1.5 mm into the tissue if an array of 3 mm in diameter is used. Such interference can be avoided using sequential activation of the electrodes; however, it does not seem to be a practical solution for the high resolution implants. Pulses of 1 ms in duration applied at repetition rate of 25 Hz (video rate) provide a duty cycle of 1/40. Thus not more than 40 electrodes can be activated without a temporal overlap. Thus, for an implant containing thousands of electrodes, sequential activation is not possible. More practical solution of the cross-talk problem is having the coaxial return electrodes around each pixel in the array. These return electrodes will localize the fields produced by different electrodes and thus isolate their effects from one another, and allow for simultaneous activation. For this reason, in the discussion below, we consider an array where each microelectrode is surrounded by the return electrode.

Voltage and Current Required for Cell Stimulation

We will use a simple passive model of extracellular stimulation [18] of the neural cell, which allows for analytical solutions, and thus markedly simplifies assessment of various geometrical and electrochemical effects.

The typical resting potential of a neural cell in the range from −60 to −70 mV and depolarization of the cell membrane by a few mV are sufficient to affect ion channels and thus elicit a physiological response, including graded potential neurons such as bipolar retinal cells [19, 20]. With extracellular stimulation a change in a cross-membrane potential is induced by application of an electric field to the surrounding medium. Since the impedance of the cellular membrane is much higher than that of the cellular cytoplasm, the interior of a cell quickly polarizes in the external electric field and its cytoplasm equipotentializes [18], as shown in Figure 14.1. Thus the cross-cellular potential actually charges the cellular membrane on two poles of the cell: polarizing the anode side and depolarizing the cathode size. In a uniform electric field these potential steps are

equal, i.e. Umembr = Ucell/2, but in a non-uniform field, which is a case with small electrode in front of a large cell, only a small area of the cell membrane is

affected thus producing much larger cross-membrane voltage drop on a proximal

end than on a remote end of the cell, so at the proximal end Umembr ≈ Ucell. The polarization time constant tp = 0 5L · cmem gcyt + g/2 , where cmem ≈ 1 F/cm2 is the membrane capacitance per unit area [21], gcyt is the cytoplasm resistivity