In the third part of this chapter we describe the dependence of the damage threshold on pulse duration, electrode size and its separation from the cells, as well as on the number of pulses. We also compare damage and stimulation thresholds to assess the safe therapeutic range at various pulse durations.

7.2  Power and Data Transmission

7.2.1  Wireline Connection

William Dobelle led one of the earliest attempts at constructing a visual prosthesis. In a series of studies begun in 1968, he used direct wireline connections to link electrodes placed in the visual cortex with a stimulator worn externally to the body [16]. Subsequent electrical stimulation successfully evoked visual responses in 19 blind patients, offering hope that future prostheses would one day restore some degree of useful sight.

Direct percutaneous connections are far from ideal, as they can provide pathogens with a direct pathway through the skin and are prone to severe scarring [30]. Despite this, transdermal cables have often been used in short-term human trials of various visual prostheses [51, 53, 65, 76], because of the unrivalled electrical versatility which they offer. For example, a group at the Naval Research Lab has developed a 3,200 electrode epiretinal prosthesis which is driven with a cable containing ten wires [58]. This prosthesis is intended for acute experiments; a future version under development is wireless. In at least one case, percutaneous cables driving a retinal prosthesis have been left in place for a period exceeding 1 year [76]. Though direct connections will likely continue to be used in research settings for years to come, any future commercial prosthesis will be wireless.

7.2.2  Inductive Coils

Inductively coupled coils are used for wireless data and power transmission in a wide variety of applications, including medical implants such as cardiac pacemakers [4] and cochlear prostheses [74]. More recently, the unique power and data requirements of visual prostheses have spurred much research in the field, with inductive coil systems currently developed for epiretinal [33, 69], subretinal [35, 63], visual cortex [64], and optic nerve stimulators [65].

In all of these designs, an AC current driven through an external transmitting coil induces an AC voltage on an implanted coil, which is converted to DC power by implanted circuitry. Sometimes the transmitter encodes data onto this signal, which is also recovered by the implanted circuitry. Since the coils are only weakly coupled to each other (typical values for coupling coefficient k are in the range 0.08–0.24 [68], compared to ~0.9 for standard transformers), great care must be taken to

optimize the receiving circuitry. With this in mind, a capacitor is added in series with the receiving coil to create a tuned resonance at the transmitter frequency, ft. The resulting circuit amplifies the received voltage by the quality factor Q, typically in the range 10–100. High Q values yield more efficient power transfer thus helping to decrease the body’s exposure to radiation. The optimization of coil geometry and receiving circuitry to maximize Q has been the subject of numerous studies [19, 20, 27, 28, 31, 62, 66, 68, 72]. Since Q is proportional to the transmission frequency, high frequency operation yields higher Q values; however, tissue’s RF absorption increases exponentially beyond a few MHz [43] limiting transmission frequency ft to 1–10 MHz.

Inductive coils have been used to deliver data to visual prostheses for over half a century. In the 1960s, a team led by Giles Brindley of the Medical Research Council in London implanted an array of 80 coils beneath the pericranium of a blind patient [8]. The 80 coils were connected through separate rectifying circuits [7] to 80 platinum electrodes placed onto the surface of the patient’s visual cortex. Individual electrodes were activated by placing a transmitting coil on the scalp directly above the electrode’s receiver. Interference was minimized by tuning adjacent receivers to different frequencies. Though this scheme was rather successful (of the 80 electrode placements, 39 elicited phosphenes), it is hardly scalable. With the goal of scaling visual prostheses to hundreds and eventually thousands of pixels, higher data rates must be extracted from fewer coils.

Ironically, while high-Q coils are efficient power receivers, they are rather poor data receivers. According to the Shannon–Hartley theorem [59], the data capacity C of a coil may be expressed as

where C is in bits per second, B is the bandwidth of the receiving circuit, ft is the transmission frequency in Hz, and SNR is the signal to noise power ratio. Thus, while received power is directly proportional to coil Q, the attainable data rate is inversely proportional to Q. For this reason, many visual prosthesis designs use two coil pairs: one for power, and one for data, where data transmission is accomplished at a higher frequency [27, 68] or with a lower-Q coil [35]. In addition, complex single-coil systems capable of delivering both power and data over one coil pair have also been developed [18, 19, 67], in one case achieving a data rate in excess of 1 Mb/s [33].

Ignoring the time involved in implant monitoring feedback signals, transmitting control signals, and other housekeeping functions, the maximum number of pixels N that can be individually driven at the refresh rate, R, is determined by the data rate, C, and the number of the stimulation strength levels S, as

For example, the system presented in [33] with a data rate of 1 Mb/s, refresh rate of 60 Hz, and 16 different stimulation levels can adequately support more than 4,096 pixels (a 64 by 64 array).

Alternatively, James Weiland’s group in USC is developing an implantable intraocular video camera [69]. Such a design would do away with the need for inductive data transmission altogether, and will require only power and (low bandwidth) control signals.

Power and data transmission efficiency are inextricably intertwined with the physical coil placements. As such, the ideal surgical placement for receiving coils is a matter of ongoing debate. In Second Sight’s first generation prosthesis, two coils implanted subcutaneously behind the ear were coupled to a second pair, attached to the frame of eyeglasses worn by the patient [23]. This location was chosen for two reasons: first, because surgeons have years of experience implanting coils behind the ear, due to the success of cochlear implants. Secondly, the outer coil may be placed very close to the implanted one, thereby maximizing coupling efficiency. The disadvantages are also twofold: first, any operation implanting both a retinal stimulator and a posterior auricular coil-set necessitates both a retinal surgeon and an otolaryngologist. This increases both the cost and the length of the operation. Secondly, a trans-scleral cable must connect the coil to the intraocular stimulator. This wire must be thin and flexible enough to allow normal eye movement, while also robust enough to withstand years of bending without failure. Fabricating such a wire is a challenge, though not an insurmountable one. For these reasons Second Sight has changed to a periocular design for their second generation prosthesis (Argus II), with receiving coils mounted on the front of the eye globe. The transmitting coils are mounted on the front of the patient’s eyeglasses [2].

The Boston Retinal Implant Project has designed a pair of coils and receiving circuitry which is sutured to the side of the eye under the conjunctiva, facing towards a transmitting coil mounted on the side arm of eyeglasses worn by the patient [63]. The German EPIRET consortium attempted a different approach, removing the lens and placing the receiving coil in the lens capsule [40]. This technique is appealing due to its similarity to cataract surgery. However, the relatively small size of the lens capsule puts tight constraints on the coil size (approximately 12 mm).

7.2.3  Serial Optical Telemetry

Photodiodes are excellent data receivers. Standard CMOS integrated photodiodes can have bandwidths in excess of 1 GHz [49], over two orders of magnitude higher than those attainable with inductive coils, whose bandwidths cannot exceed ~10 MHz due to transmission frequency and Q-factor constraints (see (7.1)). The German company IMI Technologies is developing an epiretinal prosthesis with a subconjunctivally

placed inductive coil and a photodiode placed inside the vitreous cavity [69]. The coil receives power inductively, while the photodiode receives stimulation data optically. As in the case of inductive coil systems, the data is delivered serially over one channel, so it must be decoded and distributed over the electrode array using a data processing chip connected to the retinal array by intraocular cable.

7.2.4  Photodiode Array-Based Prostheses

The Chow brothers were the first to propose and investigate the use of photodiodes in a retinal prosthesis, in the early 1990s [12]. Optobionics, a company they founded, developed a two dimensional array of 5,000 photodiodes on a 2 mm silicon disk. These so-called artificial silicon retinas (ASRs) were fabricated such that all 5,000 photodiode anodes were connected together, while their cathodes were electrically isolated from each other [48]. Each cathode contained its own iridium oxide electrode, resulting in a 2 mm device with 5,000 light-controllable electrical sources. Since each photodiode in the array collected light simultaneously, visual data could be delivered to all 5,000 pixels in parallel. This is in contrast to the serial RF and optical telemetry systems described above.

The Optobionics design assumed that ambient light would be directly converted by the photodiodes to currents strong enough to stimulate surviving retinal neurons. Indeed, initial human trials did result in some vision improvement [13]; however, this improvement was not due to electrically-elicited action potentials, but from neurotrophic effects resulting from ASR implantation [46]. Unfortunately, ambient light intensities provide insufficient current to directly stimulate nerve tissue, by at least three orders of magnitude [45]. In addition, since the electrode–electrolyte interface is capacitive when driven in a biologically compatible way, a photodiode array should be driven by pulsed, rather than continuous illumination. Indeed, later in vivo experiments on RCS rats with subretinally implanted ASRs successfully demonstrated neural activity in the superior colliculus in response to intense pulsed infrared illumination of the retina [15].

Eberhart Zrenner and his group from the University of Tuebingen, Germany have constructed a photosensitive array equipped with built-in differential amplifiers, called the microphotodiode array prosthesis (MPDA) [77]. Each pixel contains a photodiode and active circuitry which measures the difference between local and global brightness, and then drives a current corresponding to this difference through an associated microelectrode. The system is capable of utilizing almost all the electrochemical capacity of the electrodes; however, it requires separate power delivery to drive the active circuitry. So far, short-term human trials have relied on percutaneous wireline connections to deliver this power; future, wireless implants will use an inductive coil system for this purpose [6]. Recent result with a human patient confirmed the ability of the 1,500 pixel array to provide visual acuity on the order of 20/1,000, allowing a patient to read large fonts [53].

Daniel Palanker’s group at Stanford University has taken a different approach to an active photodiode array system by using video goggles to project pulsed