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

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Figure 16.2. (a) An array of 16 iridium intracortical microelectrodes extending from an epoxy superstructure. (b) A scanning electron micrograph of one of the iridium microelectrodes. The Parylene-C insulation has been ablated from the tip region, revealing the iridium metal which constitutes the microelectrode’s active surface. (c) An inserter tool for implanting the microstimulating arrays. Prior to being inserted into the brain, the microelectrodes are protected within the tool’s barrel. (d) A microelectrode array immediately after implantation into the cerebral cortex of a cat.

in the cable from dislodging the newly implanted array from the brain. This feature has proved to be especially useful when implanting multiple arrays. In the version shown, the microelectrodes are spaced approximately 380 m apart. The individual microelectrodes (Figure 16.2b) are formed from pure iridium wire by electrolytic etching. The shafts then are insulated with approximately 3 m of Parylene-C, a biocompatible insulating material, which then is ablated from the electrode’s tip region using a finely focused ultraviolet laser. This allows good control of the amount of exposed metal, which constitutes the electrode’s working surface. For our studies of microstimulation in the cat cochlear nucleus and cerebral cortex, and in the visual cortex of the rhesus monkey, we have used electrodes with surface areas of 0 001–0 002 mm2, or 1000–2000 m2 [5, 7–10]. The tips of these microelectrodes are quite blunt with radii of curvature of 4–5 × 10−6 mm (4–5 m). This blunted shape promotes a more even distribution

of the stimulus current over the electrode’s working surface [11] and it also appears to minimize the risk of injury to the microvessels of the brain [8, 12].

The electrode shafts are composed of iridium, a metal that is especially useful for microelectrodes that must function throughout the life of the patient. An iridium electrode’s working surface can be oxidized (“activated”) to form a hydrous oxide film. Then, during the microstimulation, this oxide can shuttle between valence states, and thereforeefficiently transfer charge from the electrode into the brain [13, 14]. Another important property of activated iridium is its extremely low rate of dissolution during stimulation, even when the electrode is operated at moderately high current densities; an important property of a prosthesis that must function for the life of the patient.

Figure 16.2c shows an inserter tool that we have developed for implanting microstimulating arrays into the human cochlear nucleus, but which also is suitable for implanting the arrays into the cerebral cortex [5, 7, 10]. Prior to being inserted into the brain, the microelectrodes are protected within the tool’s barrel so that the tip of the barrel can be placed in contact with the brain at the intended point of implantation. The electrodes then are injected into the brain at a moderately high velocity (approximately 1.5 m/sec). The bend at the end of the barrel facilitates access to brain surfaces that are partially obstructed by juxtaposing structures. Figure 16.2d shows an array immediately after implantation into the cerebral cortex of a cat. We also have used this tool to implant multiple microelectrode arrays into the visual cortex of 3 rhesus monkeys [7, 10]. The tool can be constructed so that the end of barrel is angled at up to about 70 (In Figure 16.2b, the angle is approximately 40 ). This feature will be valuable when implanting the microelectrode arrays into the portion of the human visual cortex that lies within the central sulcus near the brain’s occipital pole, between the two cerebral hemispheres.

Other technologies have been developed that embody the principle of the “floating” array of multiple penetrating microelectrodes, most notably the Utah Intracortical Array of 100 microelectrodes [15, 16], shown in Figure 16.3. The Utah array has been used primarily for long-term recording for cortical neurons [17], but with appropriate modifications, it could be the technology of choice for a cortical visual prosthesis, particularly if, or where, a high spatial density of electrodes is desirable [18].

Tissues Responses to Chronically

Implanted Microelectrodes

With any neural prosthesis, the risk of injury to the tissue at the site of implantation must be carefully evaluated. In the case of the intracortical electrode array, there is an opportunity for tissue injury during implantation as the microelectrodes penetrates down into the highly vascular brain tissue, and also during the subsequent period of residence in the brain. There also is the opportunity for injury from the electrical stimulation itself. The micrograph in Figure 16.4a

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Figure 16.3. The Utah Intracortical array. It contains 100 microelectrodes, and is an example of a technology for efficiently fabricating high-density microelectrode arrays.

shows the footprint in a cat’s cerebral cortex left by the 18 iridium shafts of an array of the type depicted in Figure 16.2a. The array had been implanted for 30 days. The histologic section was cut perpendicular to the axis of the electrode shafts at a depth of approximately 0.5 mm below the array’s superstructure and below the surface of the brain, and shows the tracks of the 16 working electrodes

(T) and those of the 3 longer stabilizing shafts (S).

The state of the neural tissue shown in Figure 16.4a is typical of our findings at the implant sites of these arrays. In spite of the numerous small blood vessels permeating the tissue (some of which appear similar to the electrode tracks), we have rarely seen evidence of vascular injury within the array’s footprint. We presume that as the electrodes are being inserted into the brain, their blunt tips push the vessels aside rather than severing them. Figure 16.4b shows a histologic section through the tip site of one of the microelectrodes from the same array that had been subjected to 8 hours of electrical stimulation. Figure 16.4c shows the tip site of an unpulsed microelectrode. The particulars of the stimulation regimen are described below. A conspicuous feature of the stimulated site is the aggregate of inflammatory cells (seen as irregular elongated profiles) around the tip site [19].

In Figures 16.4b, c, there are normal-appearing neurons surrounding all of the pulsed and the unpulsed tip sites. However, prolonged electrical stimulation does convey a risk of injury to nearby neurons. There are several mechanisms by which electrical stimulation might inflict tissue injury. A detailed discussion of this topic is beyond the scope of this chapter, but the interested reader is referred to the review by the author [20]. The propensity for neural damage is affected strongly by the interaction of two variables, the stimulus charge per phase and the stimulus charge density. In most cases, the stimulus waveform is configured so that the positive and negative phases inject equal amounts of charge. Charge per phase is the charge injected by the electrode during each presentation of either the positive or the negative (anodic or cathodic) phases of the stimulus current.

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Figure 16.4. (a). The footprint in a cat’s cerebral cortex left by the 18 iridium shafts of an array of the type depicted in Figure 15.2A, implanted for 30 days. (b) A histologic section through the tip site of one of the microelectrodes that have been subjected to 8 hours of electrical stimulation at 004 C/ph, and at a charge density of 200 C/cm2.

(c) A histologic section through the tips site of an unpulsed microelectrode.

A stimulus waveform that is often used for neural stimulation is composed of a pair of cathodic and anodic current pulses of constant amplitude and so charge per phase is simply I × d, where “I” is the stimulus current pulse amplitude and “d” is the duration of the pulse comprising the anodic or cathodic phase of the stimulus. The unit of electric charge is the Coulomb (C) which corresponds to 1 ampere of current for 1 second. For microstimulation, the most common units are microcoulombs 10−6 C , and nanocoulombs 10−9 C . Charge density per phase is charge per phase divided by the electrode’s surface area. It usually is expressed as microcoulombs per square cm of electrode Figure 16.5 shows how charge density and charge shows how charge density and charge per phase interact to determine the risk of neuronal injury. These data are drawn from several studies [21–24] but all employed electrodes implanted on the cat’s cerebral cortex, or microelectrodes implanted within the cortex. Filled and open symbols represent combinations of charge density and charge per phase that did or did not cause injury to nearby neurons. It is apparent that charge density and charge per phase interact synergistically to determine the propensity for neural injury.

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Figure 16.5. The relation between charge density, charge per phase and the induction of injury to neurons. The diagonal line indicates the approximate demarcation between the domains of damaging and the non-damaging stimulation.

Penetrating microelectrodes, with their small surface areas, are represented in the extreme upper left portion of the graph, and in this region charge density can be relatively high without injuring the nearby neurons. However, injury will occur if the combination of charge density and charge per phase is excessively high. Figure 16.6 shows a histologic section through the tip site of a microelectrode that underwent 8 hours of pulsing at 0.06 C/phase and at a charge density of approximately 3000 C/cm2. There are no recognizable neurons within about 70 m of the tip site. Fortunately, these stimulus parameters are far in excess of what would be required for a visual prosthesis employing ICMS. The microelectrodes depicted in Figure 16.4 were evaluated with a regimen that is more appropriate for

Figure 16.6. A histologic section through the tip site of a microelectrode that have been subjected to 8 hours of electrical stimulation at 06 C/ph, and at a charge density of 3000 C/cm2. There is marked infiltration of inflammatory cells into the tissue surrounding of the tip site (T), and no recognizable neurons near the tip.

ICMS. Eleven of the 16 microelectrodes underwent 8 hours pulsing at a charge density of .004 C/ph, and at a charge density of 200 C/cm2. This continuous stimulus was delivered at a rate of 50 pulses per second (50 Hz). Two days after the end of stimulation, the cat was deeply anesthetized and the brain tissue was examined for evidence of tissue injury. All histologically normal neurons within 100 m of each of the electrode tip sites were counted in three adjacent histologic sections through each of the tip sites, and the distance of each neuron from the center of the tip site was recorded. In this animal, the average density of normalappearing neurons around the pulsed and unpulsed electrodes (Figures 16.4b, c) differed by less than 5% between 30 and 100 m from the center of the tip site.

The selection of stimulus parameters must be constrained by the imperative of not causing neural injury. However, additional constraints are imposed by the requirement that the stimulation does not markedly perturb the physiological properties of the neurons to be stimulated, and in particular, that the stimulation does not induce large changes in their electrical excitability. If a cortical visual prosthesis is to restore useful vision to a blind person, numerous microelectrodes will have to be pulsed either simultaneously or in an appropriate temporal sequence that is yet to be determined. Prolonged neural stimulation may induce a persisting reduction of the electrical excitability of neurons that are close to the electrodes. In the cerebral cortex, this phenomena of stimulationinduced depression of electrical excitability, or “SIDNE”, [6, 9, 12] appears to be exacerbated when many closely spaced microelectrodes are pulsed [9]. SIDNE can occur in the absence of histologically detectable neural injury [12], but it remains an issue in the design of stimulus protocols since recovery of neuronal excitability after cessation of the stimulation typically requires many days, and

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when SIDNE is present, the functionality of the prosthesis obviously will be degraded and/or unstable.

We have observed persisting but ultimately reversible SIDNE in the feline cochlear nucleus and in the cerebral cortex after prolonged microstimulation [6, 9, 12], so it may be a general manifestation of the response of neurons to highly localized protracted stimulation. We investigated SIDNE in the sensorimotor cortex of the cat [9], a site which, while not identical to the primate visual cortex, is a convenient model because many of its neurons project into the corticospinal tract (the “pyramidal tract”), a bundle of axons that traverses from the cortex into the spinal cord and is sufficiently compact to be accessible with a single recording microelectrode. Arrays of 16 iridium microelectrodes similar to Figure 16.2a were implanted chronically into the sensorimotor cortex of adult cats for at least 40 days, and a recording electrode was implanted into the pyramidal tract (Figure 16.7a). Neuronal responses characteristic of single pyramidal tract axons (“unit-like responses” or ULRs) were recorded during ICMS (Figure 16.7b). Each trace was generated by averaging the response to 2048 consecutive intracortical pulses. The negative peak of the ULRs is indicated by . The graph’s abscissa is the latency after the start of the 150 s/phase biphasic stimulus pulse. The number near the right edge of each averaged trace signifies the amplitude of the 150 s intracortical stimulus pulse. The threshold of this ULR is 8 A. When the intracortical microelectrodes were not pulsed except to determine the electrical threshold of the ULRs, the electrical threshold of the ULRs was very stable for at least 7 hours (Figure 16.7c). Note that the threshold of most of the ULRs was below 12 A (1.8 nC, with the 150 s stimulus pulses used in the study). The abscissa is the ULR’s initial threshold and the ordinate is the threshold after 7 hours, during which the microelectrodes were not pulsed. The broken lines represents a change of one stimulus level, and is the limit of accuracy for the determination of the unit’s threshold. The numbers adjacent to some circles indicates multiple ULRs at those coordinates. As discussed above, eight hours of continuous pulsing of the intracortical microelectrodes at 4 nC/ph and 50 pulses per second did not induce histologically detectable neural damage, and when these stimulus parameters were applied for 7 hours to only 1 of the 16 microelectrodes in the intracortical array, there was elevation of the electrical threshold of only 1 of 18 ULRs evoked from these pulsed microelectrodes (Figure 16.7d). The data were acquired from 5 cats in which only the single intracortical microelectrode from which the ULRs were evoked was pulsed continuously for 7 hours, at 26 5 A (4 nC/ph) and at 50 Hz. The threshold of all but one of the ULRs was unchanged. However, when all 16 microelectrodes were pulsed for 7 hours at 4 nC/ph, the threshold of most of the ULRs became markedly elevated (Figures 16.8a, b), and the SIDNE was perhaps even more severe when the 16 microelectrodes were pulsed sequentially (at 50 Hz per microelectrode) (compare Figures 16.8a and 16.8b).

We postulate that reversible SIDNE is caused by the prolonged high-rate neuronal activity that is induced by the stimulation, and when the stimulus amplitude is sufficiently high so as to produce significant overlap in the regions

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Figure 16.7. (a) The scheme used to study stimulation-induced change in neuronal excitability in the cat’s cerebral cortex. The array of stimulating microelectrodes is implanted into the sensorimotor cortex and a single recording electrode is implanted though the cerebellum and into the pyramidal tract. (b) A “unit-like response” (ULR) evoked by microstimulation in a cat’s postcruciate gyrus and recorded from the pyramidal tract. (c) The thresholds of 26 ULRs (circles) from 2 cats. (d) The effects of 7 hours of continuous stimulation on the threshold of 18 ULRs (From [9]McCreery et al, 2002, with permission of the publisher).

of effective stimulation from adjacent microelectrodes, sequential stimulation will evoke multiple responses from neurons in this overlapping region, forcing the neurons to fire at a very high rate. When the stimulus amplitude was lower (e.g. 1.8 nC/ph, Figures 16.8c, d), then sequential pulsing of the 16 microelectrodes (Figure 16.8c) induced much less SIDNE than simultaneous pulsing of the 16 electrodes (Figure 16.8d).

those particular neutons. However, the fact that 7 hours of simultaneous pulsing at 1.8 nC caused a greater effect on the thresholds of these low-threshold ULRs suggests that the effects of the prolong stimulation on neuronal excitability are determined not only by the forced activation of the individual neurons, but also by the total number of neurons activated. Simultaneous pulsing of adjacent microelectrodes will induce neuronal activity over a greater distance than will sequential pulsing, due to summation of the stimulus from each electrode. While the quantitative details certainly will differ for microelectrode arrays with different interelectrode spacing, or when pulsed with different parameters, or when implanted into different regions of the cerebral cortex (visual vs. somatosensory cortex), and perhaps even when implanted into different species (humans vs. domestic cat), it is likely that SIDNE in a cortical visual prosthesis could be minimized by using a stimulation protocol that minimizes interaction between the microelectrodes (and thus also maximizing the spatial resolution). Sequential pulsing of the microelectrodes at an amplitude for which there is minimum overlap of the effective stimulus from each electrode would meet these requirements.

While SIDNE obviously will be an issue in the selection of the stimulus parameters to be used with a visual prosthesis employing ICMS, it is gratifying that we were able to deliver sequential ICMS for many hours through many adjacent microelectrode at an amplitude of 1.8 nC/ph and 12 A, without inflicting histologically detectable neural injury and without inducing detectable SIDNE in most of the low-threshold neurons near the microelectrodes. This stimulus lies within the range that was shown to induce phosphenes in a human volunteer with microelectrodes implanted in the visual cortex (1.9–25 A; [1]) which produced directed saccadic eye movements when delivered into the deep layers of the primary visual cortex of a rhesus monkey (10 A or less; [2, 25]) and which apparently is appropriate to convey a percept of motion to an object in the visual field when delivered via a microelectrode in cortical area MT (10 A; [3]).

In our laboratory, microelectrode arrays similar to the example shown in Figure 16.2a have been implanted into the cat cerebral cortex for up to 2 years [5], and have remained functional over that interval. Most of our experience with microstimulating arrays implanted for long intervals is derived from our program to develop an auditory prosthesis for person with severe hearing loss, and who are unable to benefit from a cochlear implant. We have implanted more than 50 arrays into the cats’ cochlear nucleus, for up to 7 years [8, 12, 26, 27]. The cochlear nucleus receives its input from the inner ear via the auditory nerve, and is the first stage of processing of auditory information in the central nervous system. These animal studies have yielded valuable insights into the histologic and physiologic response of the neural elements surrounding penetrating microelectrodes that have been implanted for long intervals. Figure 16.9a shows a histologic section through the site of the tip of one of four microelectrodes that had been implanted in a cat’s cochlear nucleus for 2588 days [8]. The histologic section was cut through the microelectrode’s track at an oblique angle, so the microelectrode’s tip site appears elongated. The tips site is encapsulated with a sheath of connective tissue approximately 20 m in thickness. Beyond this capsule, the fibrous network