Ординатура / Офтальмология / Английские материалы / Visual Prosthesis and Ophthalmic Devices New Hope in Sight_Rizzo, Tombran-Tink, Barnstable_2007

Fig. 2. (A) Diagram of the electrode-cell geometry for Eq. (1). (B) Increasing distance from Y1 = 5 to 100 m increases the required threshold current density at the electrode. Threshold current is the current required to create a ∆V equal to 30 mV. (C) The effect of distance on resolution. Each cell (circles) experiences a ∆V from each active electrode (dotted arrows). These ∆V are the result of the electric field from each electrode and thus, add vectorally to result in a total ∆Vk (black arrows) for each cell. The center cell in the case of Y1 experiences a ∆V1, which

is lesser than ∆V2,3, thus the retina resolves the two discrete electrodes. This is not so for the more distant case of Y2, here the ∆V1 is about the same as ∆V2,3. (See text for details.)

Figures 2B,C present the two effects of distance, increased current density at the electrode surface and loss of resolution. Figure 2B presents an example of the increased current density required to maintain the desired ∆V = 30 mV. Using Eq. (1), for a 20m diameter electrode a 95 m increase in distance, from Y = 5 to 100 m, will cause the required current density to increase by a factor of 35, from 1.6 to 56.6 A/cm2. The resolution effect is based on the fact that the voltage difference, ∆V, across a cell is dependant on all the current sources in the vicinity of a given cell. In Fig. 2C, for the two active electrodes, E1,2, each cell (circles) experiences a ∆V from each active electrode (dotted arrows). These ∆V are the result of the electric field from each electrode and thus, add vectorally to result in a total ∆Vk (black arrows) for each cell. The cells in each row and the electrodes are separated by a distance x, which is used in determining the vectors ∆Vk. For the three cells at a distance Y1 from electrode array, the difference between the total ∆V vector (sum of two dotted lines on the cell) of the center cell (∆V2) from those of the side cells (∆V1,3) is sufficient such that for the correct current the center cell will not be stimulated whereas the side cells are. The three cell “retina” can resolve that two separate electrodes are on.

This is not the case for the cells at Y2, the larger distance. Here the total ∆V experienced by each cell is virtually the same because the distance of side cells and the center cell from the electrodes is almost the same. The three cell “retina” perceives the two active

Fig. 3. Threshold current required to generate 30 mV drop across a 10 m long cell, plotted as a function of distance between the cell and the electrode surface. The four curves correspond to current calculated for electrode radii of 5, 15, 50, and 150 m.

electrodes as one. As an example, for x = 30 µm, Y = 5 µm, je = 1.6 A/cm2, γ = 70 W,

and L = 10 µm eq. (2) results in ∆V1,3 =0.030 V whereas ∆V2 = 0.002 V. Changing Y to 100 µm and je to 56.6A/cm2 results in ∆V1,3 =0.051 V whereas ∆V2 = 0.053 V. Of course the use of hemispherical electrodes and 10 µm cells makes a number of simplifications

concerning cell shape, axons, and dendrites; the general principle is the same for real retinas and electrodes.

A comparison of two early studies of human retinal stimulation shows that reduced distance between the electrode and the retina may reduce thresholds. In one study a gold weight fixation method was used with arrays of electrodes of 100 or 400 µm diameters (11). They found that the majority of the reported thresholds were lower than in a study using a hand held, more distant from the retina electrode (12), 0.28–2.8 mC/cm2 vs 0.16–80 mC/cm2 respectively. Even within the study by Rizzo et al. they found lower thresholds in general for a smaller, weighted 100 µm diameter electrode vs a 250 µm hand held electrode, 50 µA vs 500 µA. A seemingly contrary position is taken in a recent paper by Mahadevappa (13), suggesting that the distance from the retina is not a factor in thresholds in practice until the distance is more than 500 µm. The graph by Palanker et al. indicates that this would likely be the case for a larger 400 µm electrode studied by Mahadevappa. However, 400 µm electrodes will not result in high-resolu- tion retinal prostheses. The smaller electrodes required for a high-resolution retinal prosthesis will require close contact.

EPIRETINAL OR SUBRETINAL PLACEMENT

A primary division in retinal prostheses is between epiretinal and subretinal (14,15). The descriptive difference is the placement of the device against the inner limiting membrane, epiretinal, or between the retina and the sclera, subretinal (see Fig. 4.) The physiological differences are many. A key physiological difference is that epiretinal

Fig. 4. Diagram indicating the general subretinal and epiretinal placements of retinal prostheses. Not drawn to scale.

devices are closer to the ganglion cells. Thus, they may require preprocessing of the image before stimulation to make up for the bypassed function of bipolar, horizontal, and amacrine cells. Subretinal devices are more likely to stimulate bipolar cells because they are closer to the bipolar cells than to the ganglion cells and thus, in theory, preprocessing may not be required. Another difference is in the potential effect of the diseased retina on the imaging of the implant. In the severest form of age-related macular degeneration (AMD) that likely to lead to blindness, the retina becomes more opaque. This opaqueness would make it difficult for subretinal devices with electronic imaging arrays located behind the retina to collect sufficient light to form an image. The surgical approach is different for epiretinal vs subretinal implantation. The epiretinal approach requires a vitrectomy and entering the eye. The vitrectomy is a standard procedure for a retinal surgeon.

FLEXIBLE DESIGNS

Many of the research groups developing retinal prostheses use flat metal arrays on polyamide. Polyamide has the benefit of flexibility, if only in one direction. This allows the device to conform in one direction to the curvature of the retina. The designs of most devices using polyamide connect the polyamide stimulation array directly to the retina. The limitation of this design is that a lead is required for each electrode. This may reduce the number of electrodes possible because of space limitations and routing complications. Most systems using polyamide currently are using less than 100 electrodes. In fact, the four groups described in this chapter have fewer than 30 electrodes. Using the latest technology DRC, Inc specifies 5 m as the smallest reliable line size and spacing. Each line therefore requires 10 m. The limit for a 3 mm wide polyamide cable would be 300 lines. Far fewer than the thousands required for reading and face recognition. Layering signal lines is possible, but could create capacitive coupling problems and reduce polyamide array flexibility.

Polyamide electrodes are formed using standardized methods (Fig. 5) (16). The first step is to spin a layer of polyamide onto a solid substrate, such as a silicon wafer and cure it in an oven. In some facilities, large prefabricated sheets of polyamide are used

Fig. 5. Steps to fabricate polyamide electrode arrays and cables. Polyamide is spun onto a substrate and gold and/or other metals are patterned on (1–4). Aluminum is patterned as a mask to block the reactive ion etch of polyamide in selected areas (5,6). A thin wet etch removes the Al (7). Finally the device is lifted of the substrate (8).

instead. The thickness of the base layer can vary, but 5–15 µm is typical. The next step is to deposit a layer of metal, typically sputtered gold, to act as the interconnections and other device features. Lift-off technology is used to pattern the interconnections and features in a photoresist and remove excess gold to reveal the desired pattern. An additional metallization step to add patterned biocompatible electrode materials, such as platinum or iridium may be performed if those metals have not been used as the base metal. A second layer of polyamide is spun to provide insulation of the metal features. In some cases, before the second layer of polyamide added, a layer of silicon nitride is used to improve impedance. Sputtered and wet etched aluminum can then be used as a mask to allow the selective reactive ion etching of the polyamide to reveal windows to the electrodes and other polyamide device features.

The Tubingen, Intelligent Implants (IIP), and Boston VA groups all use variations of polyamide with metal traces leading to openings in the top layer polyamide to form the electrode sites (Fig. 6). The Tubingen group published a recent article for a device with 16 electrodes on polyamide (3,17). Their electrode is for subretinal implantation and the active area covers approx 1 × 1 mm2. The Boston VA group has had a number of designs in the past for both epiand subretinal implantation. Their current design is a subretinal array similar to the Tubingen group. They have the ability to make their electrode array from made of either polyamide or parylene-C. They report results with a test version of three device that has 25 electrodes in an 1 × 1 mm2 array (18). The IIP group also uses a flat polyamide array with exposed electrodes to connect their electronics to the retina (6). The device, as with the Tubingen and Boston VA groups has a portion of the polyamide, which acts as a cable to transmit current laterally from the electronics before directing current into the retina. The IIP device is different in that they plan to create an epiretinal device.

Two unique designs use micro electro mechanical system (MEMS) type fabrication methods to bring their arrays in close approximation with the retina (Fig. 7). A design of the Boston VA group, in collaboration with the Cornell Nanofabrication Institute and MIT, involves a flexible epiretinal array, which can unfold when inserted into the eye

Fig. 6. Three examples of polyamide arrays/cables used for prototype retinal prostheses. (A) The Tubingen group device with array attached. (B) Two test versions of the IIP device (3,6).

Fig. 7. MEMS type flexible devices with three-dimensional structures. (A) Boston VA group inflatable array with 100 electrodes supported by a cannula (1) (image courtesy of D. Shire). (B) Stanford/Palankar recessed holes design showing migrated retina cells (scale bar = 50 m) (2).

through a narrow incision. This interesting design allows for a large area of 9 mm diameter to be covered by a device that could be easier to implant because of the small 3 mm incision required to implant. In addition, inflatable channels within the device would gently press the arms of the device and the electrodes against the retina to improve contact. The Stanford/Palanker group has a different approach. They hope to bring the retina closer to the electrodes by creating chambers in a multilayer polyamide membrane. Their results from in vitro culture and in vivo 9 d implantations show that the outer retinal cells will migrate into the recesses of the subretinal device (2). Their concept for a subretinal device would require lower current levels because of the close approximation of the electrodes and the cells. Some issues remain, such as how well the cells will survive in the recesses.

Two other flexible designs use a silicon rubber, or polydimethylsiloxane (PDMS), to create retinal electrode arrays (Fig. 8). Lawrence Livermore National Laboratory has a PDMS device, which is very similar to the flat polyamide devices with gold traces

Fig. 8. Two flexible Polydimethylsiloxane devices. (A) Doheny/Second Sight 16 electrode device is a thick Polydimethylsiloxane array molded to have a curvature to fit the retina (7,8). (B) A polyamide like electrode array developed at the Lawrence Livermore National Laboratory (10).

described earlier (10). As with most polyamide designs, this type of design allows for a cylindrical curvature, but not a spherical curvature to fit the retina correctly. The thin PDMS can be distorted toward a spherical curvature, but then the electrodes will not have regular contact with the retina because of folds. A difference from polyamide devices is the addition of a thicker PDMS “rib,” which extends along the cable portion of the device for support.

Second, Sight, Inc has a device made of silicon rubber, which is currently under clinical testing by USC/Doheny Eye Institute (13). This device has 16 electrodes, 500 µm in diameter embedded in the silicon rubber. There is a cable made of silicon rubber and continuous with the electrode pad, which contains the wire leading to the stimulation current source. Unlike flat polyamide or PDMS devices the silicon is molded into a spherical geometry to fit the surface of the retina and make good contact to reduce thresholds. However, while flexible the PDMS array is thick enough that it may not readily conform to the retina. The difficulty in this design is how to scale up to hundreds or thousands of electrodes with a method that does not use silicon fabrication methods or MEMS techniques.

MEMS/HYBRIDE DESIGN

Sandia National Laboratory has a unique design that attempts to meet the challenge of adapting to the spherical curvature of the retina (4,5). Their design uses MEMS methods to create springs for electrode posts “float” on, Fig. 9. The springs are made from micromachined silicon. In the device reported in their 2005 conference paper (19), they detail implantation of 5 × 5 mm2 array with 81 electrodes. They reported a few minor problems with inserting the device into the scleral incision, getting all the electrodes in contact with the retina and damage to the electrode posts. They solved some problems with an insertion sleeve and continue to improve their device. The obvious advantage is the ability in theory to match the curvature of any retina if the difference

Fig. 9. SEM image of the Sandia device (A) and a cross-sectional diagram of the same (B). In the SEM picture, the white bar is 500 m. The center-to-center spacing of the electrodes is approx 500 m (4,5).

in surface elevation is less than 100 m, the maximum deflection distance of the springs. A high-resolution device would require electrodes with a separation of 50 m or less. The separation between their electrodes is 500 m. As designed, this could be problematic, as higher resolutions will require closer spring spacing. The geometry of their device indicates that the deflection distance is proportional to the spring width and closer spacing would likely lead to reduced deflection distance. However, as the device uses standard MEMS fabrication methods manufacturing higher resolution devices is possible in theory.

RIGID DESIGNS

One of the earliest rigid designs is that of Optobionics, Inc. (Fig. 10). The prosthesis is a subretinal 25 m thick silicon chip with 5000 photodiode electrodes. The device is flat and nonconforming to the shape of the retina. However, as the device is implanted subretinally the retina can be pushed toward a flat profile. The device is 2 mm in diameter, which results in a 30 m distance from the edges of the device to the retina. This distance may be small enough that the increase in current and loss of resolution may be minimal. The Optobionics (Naperville, IL) evice suggests that smaller devices may be way to avoid conformal fit to the shape of the retina. Of course, the field of view for such a small device is limited.

Two groups working at Stanford have produced rigid devices with protruding electrode columns. An interesting rigid design, developed by the Stanford/Harris group (20), involves an array of carbon nanotube (CNT) columns. The columns are grown on a rigid silicon substrate. The CNT electrode array is designed to penetrate the retina to allow for close contacted between the retinal neurons and the electrodes. The CNT columns have some degree of flexibility, which allows them to cause less damage as they penetrate the retina. The design makes some accommodation to the curvature of the retina in that electrodes at the edges of the device could penetrate slightly into the retina to allow center electrodes to get closer to the retina. The columns have been grown to heights of 100 m, at a height:width aspect ratio of 4:1.

Fig. 10. The Optobionics retinal prosthesis. The 2 mm silicon disk (A) is placed behind the retina (B).

Fig. 11. Rigid columnar retinal prosthesis electrodes. The Stanford/Palanker group lithographically fabricated silicon columns 10 m wide and 70 m long penetrate from the photoreceptor cells to the inner nuclear layer (2).

The Stanford/Palanker group has a concept for a pillared electrode array made on a lithographically fabricated silicon wafer (Fig. 11). The device would be 3 mm in diameter. They implanted a prototype device with 10 m columns in a RCS rat for 15 d. Their experiment demonstrated the ability of the columns to penetrate (Fig. 11C).

As with the Optobionics device, the Naval Research Laboratory (NRL) device is a rigid device based on a silicon chip (21–24). The 3200 electrode, NRL device is 6 mm long and 3 mm wide. The resulting gap between the silicon chip and the retina would be as great as 360 m at the center. Unlike the Optobionics device, the NRL device is hybridized through indium bump bonds to an array of microwires imbedded in a glass matrix. In addition, the NRL device is an epiretinal device, which requires a conformal fit or a significant space between the center of the device and the retina will result. The microwire/glass array acts as a bridge material to span the gap between the flat silicon chip and the retina. It conducts the stimulus current from electrode to the retina through the microwires, but it insulates one electrode from the others with the glass matrix material.

Microchannel glass is fabricated using glass-drawing procedures that involve bundled stacks of composite glass fibers (Fig. 12 B,C). An acetic acid-etchable glass rod is inserted into a nonacetic acid-etchable glass tube. This pairing of dissimilar glasses is drawn at an elevated temperature into a fiber of smaller diameter. Several thousand of these fibers are

Fig. 12. The Naval Research Laboratory stimulation array. (A) The side view of a partially encapsulated device showing the curvature of the device surface. (B) A high magnification SEM of the 5.5 m microwires (bar = 10 m). (C) A low magnification SEM of the side of the microwire glass. A few microwire can be seen to be bent from the glass cleaving process (bar = 100 m). (D) the completed retinal prosthesis. The surgical sponge on the polyamide cable is used to reduce leakage of saline during the acute implant procedure.

then cut and stacked in a hexagonal-close-packed arrangement, yielding a hexagonalshaped bundle. This bundle is subsequently drawn at an elevated temperature, fusing the individual composite fibers together whereas reducing the overall bundle size. At this stage, the fibers are hexagonal-shaped and contain a fine structure of several 1000 m sized (5–10 m diameter) acid-etchable glass fibers in a hexagonal-close-packed pattern. These fibers are then bundled together in a 12-sided bundle and fused together at an elevated temperature. Standard microchannel plate glass is obtained from the combined boule at this point by slicing thin 200–1000 m thick wafers, which are polished flat. Wafers are then placed in acetic acid to remove the acid-etchable glass. In this way, a glass with extremely uniform, parallel channels is obtained.

Next, electroplating is performed to fill the channels with a metal to create microwires. The high, approx 200:1, aspect ratio of the channel length to channel diameter makes it difficult to electroplate the metal. Thus, the channel filling metal must readily electroplate. One such metal is nickel, which is easier to electroplate than gold or platinum. As nickel is highly electro-chemically active, it is not biocompatible as an electrode. However, this has been found that it is possible to “cap” the nickel microwires with a layer of gold, platinum, or other suitable electrode material.

Microwire glass electrodes used for eventual in vivo testing with the stimulation array will have one side of the electrode curved to create a spherical surface to allow positioning of the high-density electrode array in extremely close approximation to the retinal tissue. The radius of curvature is nominally 12.7 mm to provide a conformal fit against

the retina. In the future, the NRL device could be custom made with diamond turning or other polishing methods standard to optics industry to match any curvature of the patient’s retina as measured by a profilometry device such as optical computed tomography.

The polishing process will create slightly recessed microwires with respect to the curved microwire glass surface because the metal is softer than the glass. Therefore, further processing is necessary to create electrodes that protrude slightly above the curved surface. This can be accomplished by applying a 5% HF/5% HN3 etch to the surface that removes several microns of glass. The duration of this etching step can be increased to provide longer protrusions of microwires, which reduce the electrode impedance by increasing the metal microwire surface area in contact with the saline environment of the eye. A scanning electron micrograph of microwire glass having channel and microwire diameters of 5.6 m is shown in Fig. 12.

CONCLUSIONS

Almost all of the device designs presented in this chapter make some attempt to account for the curvature of the retina. Most provide for cylindrical curvature and a few provide for spherical curvature. The true advantage of curved electrode arrays is only apparent for high-resolution devices. A 300 m distance from the retina is not important for 300 m diameter electrodes. However, with the exception of the NRL and Optobionics arrays, there appear to be no fully functioning implantable devices with resolution high enough to take advantage of the curvature accommodations of the devices described in this chapter. As the Optobionics device offers little in the way of curvature accommodation and the NRL device has not been tested in humans at the time of this publication, no clear experimental evidence exists to say, which designs described in this chapter are the best. What is clear is that there exists a great diversity of approaches to meeting the challenge of the curved retina. With such diversity it is likely that an effective solution will be found and the challenge met.

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