Implant Electronics for Intraocular EpiretinalNeuro-stimulators

Torsten Lehmann∗, Nigel H. Lovell†, Gregg J. Suaning†, Philip Preston†, Yan T. Wong†, Norbert Dommel†,Louis Hyunsuk Jung†, Yashodhan Moghe∗ and Kushal Das∗

∗School of Electrical Engineering and Telecommunication, The University of New South WalesUNSW Sydney, NSW 2052, Australia, Email: tlehmann@unsw.edu.au

†Graduate School of Biomedical Engineering, The University of New South WalesUNSW Sydney, NSW 2052, Australia, Email: n.lovell@unsw.edu.au, greggs@ieee.org

Abstract—In this paper, we discuss system architectures, design chal-lenges and circuit implementation principles for intraocular epiretinalneuro-stimulators to be used as part of vision prostheses in partiallyrestoring vision to the blind. Our unique hexagonal electrode placementallows focused simultaneous stimulation which allow the electrode countto scale. A new dual-channel transcutaneous inductive link is used totransfer power and data efficiently to the implant, and a new dual-voltage power rectifier provides two implant supplies without the needfor a DC-DC converter. Low-power designs of current stimulators andECAP amplifiers are also outlined in the paper.

I. INTRODUCTION

Diseases such as retintis pigmentosa and macular degeneration, at-tack primarily the photoreceptors in the retina, and cause progressiveloss of vision [1]. The optic nerve and its extension onto the retina,the retinal ganglion cells (RGCs), however, are left relatively intactby such diseases. It is has been demonstrated in humans with diseasedretinas that nerve cells can be activated by electrical stimulation; assuch, by electrically stimulating retinal ganglion cells, neural signalstravelling down the optic nerve to the visual cortex can be generated,evoking the perception of light (a phosphene) [2]. It is thus plausiblethat a therapeutic neuro-stimulator can be implemented which canconvey a visual scene by stimulating on a number of sites on theretina via a 2-dimensional electrode array, thereby partially restoringvision in the blind. We have investigated the implementation of suchretinal neuro-stimulators for a number of years, and have developeda range of novel circuits that address the design challenges of suchvision prostheses. In this paper, we outline our prosthesis systemdesign, and system design considerations, and discuss our low-powerdesign approaches to several central implant circuits.

II. IMPLANT SYSTEM

A number of different system architectures for vision prosthesishave been proposed, including cortical stimulators, optic nerve stim-ulators, sub-retinal stimulators (where the stimulator is located behindthe photoreceptors), and epiretinal stimulators (where the stimulatoris placed closest to the retinal ganglion cell layer) [3]. Our approachis an epiretinal stimulator placed wholly within the ocular anatomyand having power and data sent to this implant via a transcutaneous(or rather a trans-scleral) inductive link [4], see Fig. 1. Such anarchitecture has a number of important properties: 1) no wires extendthrough the eye boundary, reducing the risk of infections and leavingthe eye free to move; 2) the transcutaneous link implant coil isplaced close to the front of the eye, enabling good coupling to theexternal coil, thus facilitating reasonably effective power transfer tothe implant; 3) the stimulation appears early in the visual pathwayto reduce the required pre-stimulation signal processing.

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Fig. 1. Bionic eye, comprised of an intra-ocular epiretinal neuro-stimulator(implant) and external support electronics mounted on a pair of spectacles.

group 1 electrodesgroup 2 electrodes

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Fig. 2. Concurrent stimulations on hexagonally arranged electrode array;cathodic stimulation phase with bottom-right electrodes in each group beingthe active electrodes. Groups of seven electrodes form unit cells. Arrowsindicate the direction of current flow between electrodes.

While there is a one-to-one correspondence between an imageapplied to the photo-receptors in a healthy eye and the image per-ceived, a one-to-one correspondence between phosphene location andstimulation location cannot in general be assumed. This is because ofthe neural processing that takes place on the retina, and because ofthe spread of stimulation current over many excitable neurons withcurrent electrode technology. Thus, there is a complex relationshipbetween the sites being stimulated and the image perceived bythe implantee, and one important function of early retinal neuro-stimulators is to gain a better understanding of this relationship. Itis generally accepted, though, that the perceived image quality willbe a strong function of the number of possible stimulation sites;psychophysical studies have shown that several hundreds “pixels”are required for useful vision restoration [5].

In order to address the issues of stimulation current spread andnumber of stimulation sites, we have devised a unique hexagonalarrangement of the retinal electrodes, see Fig. 2: Seven electrodes (agroup) are arranged in a hexagonal pattern which is repeated overthe electrode array. By means of a suitable switch-matrix, each ofthese can be configured as the active stimulating electrode, driven by

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Fig. 3. Concept diagram of implant stimulating tissue concurrently onmultiple electrode groups. Post-stimulation all switches are closed. For leakagedetection, amp-meters are inserted at positions marked “x”.

a stimulation current −Istim, while the six electrodes surrounding it(some possibly from other groups) act as return electrodes, driven bya total return current Istim (see [6] for details). The major advantage ofthis electrode arrangement is that the current flow in the retina tendsto be confined to just below the stimulating electrode, thus facilitatinglocalised stimulation of the RGCs and densely packed electrodes [7].By using current sources to provide the return current (Fig. 3), it isalso possible to stimulate one electrode in each group simultaneouslyas shown in the figure: with N groups of seven electrodes, all 7Nelectrodes can be employed as active stimulation electrodes in sevensequential steps (first using the bottom right electrode in each groupas the active electrodes as show in Fig. 2, then using the bottom left,and so forth). This parallelism is crucial in order to ensure flicker-free perception when the electrode count scales up to many hundreds.Our system architecture is scalable and built with blocks that eachcontrol one group of seven electrodes; the blocks consist primarilyof a current source-sink pair and multiplexing switches.When electrically stimulating body tissue, it is important to avoid

any DC currents through the electrodes as these can cause electrodecorrosion and the generation of toxic species that damage the tis-sue [8]; due to the high number of electrodes in a vision implant,using DC blocking capacitors is impractical. Instead, all electrodesare shorted to a tissue reference voltage VTR after a bi-phasic stimula-tion sequence (ideally zero net charge), to flush out any charge build-up, for instance due to mismatch in current sources (see Fig. 3) [8].This shorting also serves to define the tissue-implant potential, whichwould otherwise be undefined, ensuring sufficient voltage head-roomacross all current sources prior to the next stimulation, thus we chooseVTR � VDDH/2, where VDDH is the stimulating power supply. Notethat during shorting, the tissue is essentially a supernode with onlyone connection to the implant; hence any noise on VTR (for instancedue to noise onVDDH) cannot couple into the tissue. As all stimulatingcurrent sources cannot be identical, the implant-tissue voltage willmove up or down during stimulation causing (at least) one currentsource to go out of compliance such that the net current into the tissueis zero. While this is not a major concern as long as inter-stimulishorting is employed, the effect can be prevented by including onetissue reference electrode that can take up any current difference.

A. Design ChallengesThe design of the implant electronics in such a vision prosthesis

system poses a number of challenges compared with other electronicprostheses such as cochlear implants or pacemakers. Firstly, the num-ber of distinct stimulation sites required for useful visual perception

may be many hundreds. As a consequence, the electrode size —with current electrode technology — has to be comparatively small,increasing the electrode impedance, and hence the power requiredto deliver perceptual stimuli. Although the stimulation repetitionfrequency is low compared with that of cochlear implants, theoverall effect is an increased power requirement to the stimulatingelectrode array. Secondly, the available implant volume is restrictedto what can conveniently be attached inside the ocular anatomy.This means that there is very little opportunity to utilise off-chipcomponents, such as inductors; further, it means that the amountof power that can be safely transmitted (see [9]) to the implant ismore restricted than in other types of electronic prostheses owingto a smaller antenna diameter. Thirdly, inside the ocular anatomy,there is little heat-exchange via convection. In order to prevent anoperating implant from increasing the temperature of the eye to suchlevels where damage would occur ([10]), the overall implant powerdissipation must be kept within a few tens of milliwatts. Finally,the psychophysics of retinal neuro-stimulation is complex and notfully understood. As such, to enable clinical research into successfulstimulation strategies, the implant stimulation capabilities must bedesigned with maximum flexibility in mind, and the implant mustbe able to provide feed-back from the stimulations such as recordedelectrically evoked compound action potentials. As such, minimisingpower dissipation in all aspects of the implant design is the maindesign challenge.Another design consideration is the technology choice; smallretinal electrodes have impedances in the order of tens of kΩ, which— for stimulating currents up to 1mA — require tens of volts acrossthe stimulation electrodes. As such, a high-voltage CMOS technologyis required. In order to save power and chip real-estate, high-voltagecomponents (denoted by HV) are used sparingly, only when required.

III. CIRCUITS FOR IMPLANTSA. Transcutaneous LinkIn biomedical implants, power and data are usually sent to theimplant via a transcutaneous set of coupled inductors which form partof an LC tank tuned to the link carrier frequency. The link carrierfrequency and tank quality factor are chosen as a compromise toallow a reasonable power efficiency while achieving a sufficient datarate using ASK. Our transcutaneous link, however, is implementedusing a new dual-channel power and data inductive link in order toallow individual optimisation of power and data channels [11], seeFig. 4.The power channel uses a pair of spiral inductors that are placedon the same axis of rotational symmetry in order to maximisethe coupling between the inductors and hence the power transferefficiency. The inductors are manufactured by laser milling 35μmthick copper sheets on polyimide tape and are cascaded to allowflexibility in choice of inductance. As power rectification circuitshave often higher efficiency at lower frequencies and as the effectiveinductor series resistance increases at high frequencies, the powercarrier has been set to be below 500 kHz.The data link (externals-to-implant) uses a pair of inductors thatare placed perpendicular to the power inductors such that the netmagnetic flux from the power link is, ideally, zero. To further reducethe coupling between the two channels, frequencies above 4MHzare chosen for the data channel. The relative positions of the datachannel inductors cause the coupling between them to be ratherweak; however, as no power is transferred over the channel, eventhis weak coupling is sufficient to achieve a reliable data link.For reverse telemetry (implant-to-externals), we use a third pair of

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Fig. 4. Dual-channel inductive power and data link; front view and sideview. Magnetic flux is indicated with dashed lines. A third channel can beadded by adding a second pair of data inductors rotated 90◦ .

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Fig. 5. Dual-channel power reception. All low-voltage sections are poweredfrom VDDL to reduce power dissipation; the VDDL diode can be replaced withan active diode to further increase efficiency.

coupled inductors of similar geometry to the data link, mountedperpendicular to the first two inductor pairs.

B. Power RectifierImplant power supplies powered via a transcutaneous link are often

generated by a simple half-wave diode rectifier tapped from the LCtank in the implant side of the transcutaneous [12] link (see Fig. 5).Stimulating electrodes require fairly high voltages (some 5V-20V)for proper operation which makes the half-wave rectifier a powerefficient implementation. Most of the implant electronics, however,operate more power efficiently from a lower supply voltage (say 3Vor lower) where low-voltage transistors can be used. In an intra-ocular implant there is no volume for power converting inductors.Thus, to avoid power converters in the implant, we have deviseda novel way to extract power from the transcutaneous power linkdirectly into two (or more) different power supplies as shown inFig. 5. Power to the high power supply, VDDH is extracted as usualusing a single diode rectifier. Power to the low power supply, VDDLis likewise extracted using a single diode rectifier, but this is gatedby a transistor which conducts only if VDDL is below a set referencelevel. VDDL thus has priority over VDDH which gives correct systemoperation if the implant control system runs on VDDL.Depending on the implementation details of the VDDL power

control system, power may not be delivered to VDDH the cycle afterpower has been delivered to VDDL. In systems that do not use aseparate data link this may interfere with the implant communication.Because of the comparatively low frequency of the power channel, thecontrol system can be implemented with good efficiency. This alsoenables efficient implementations of active rectifiers such as reportedin [13] to increase the efficiency further.

C. Implant Data Receiver and TransmitterAs wireless data channels operate at relatively high frequencies

(typically above 5MHz in biomedical implants) compared with whatis required for other implant functions, the data channel can consumea significant amount of the available implant power. By using FSK, ona dedicated channel for data to the implant, only zero-crossings needto be detected by the implant receiver circuit, enabling a very simpleand power efficient implementation (see Fig. 6). Data is transmitted

clockrecovery delay sync. clock

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Fig. 6. FSK Data reception.

at 4MHz (0) and 8MHz (1) which is chosen as low as possiblewith ample frequency separation while still achieving sufficient datathroughput. The FSK receiver also extracts the implant system clockfrom the data which avoids the need for an implanted, power-hungryoscillator. For implementation details see [11].For reverse telemetry, we use simple base-band OOK. To transmita one, we excite the telemetry channel LC tank with a short pulsegenerated by an H-bridge. The LC tank is tuned above 10MHz toreduce interference with the other channels. Again such a simpleapproach enables a very low-power implementation on the implantside.

D. Stimulator DAC

Neural stimulators usually use a current to stimulate the excitabletissue as this gives the best control over the physiological responseunder varying electrode impedances. The stimulation current needsto be varied for each stimulation site according to the phosphenethreshold and the desired intensity of the phosphene. The implantbeing under digital control, a current output DAC is used to gen-erate this current. The stimulation current has normally a very lowduty cycle (for seven electrodes controlled by one current sourcestimulating on two 100μs phases at 50Hz, the duty cycle is 7%).When stimulating, the current source may draw up to about 1mAfrom the high power supply; it is thus imperative that the currentsource can be switched into a low-power, stand-by mode (currentdraw in the order of 100nA) and be switched in and out of thismode quickly (in the order of a few μs). The key to achieving a fastturn-on time while keeping the stand-by current draw low is to keepinternal capacitances small. Although the DAC resolution does notneed to be high (5 bit suffices), the stimulation and return currentneed to match well. We use a DAC where each 3 MSB thermometercoded current levels are individually trimmed at startup to achievegood cross-DAC current matching without large transistor areas, thuskeeping internal capacitances low. One MSB of such a DAC is shownin Fig. 7. Transistor areas are kept low by varying gate voltages,transistor lengths (L) and transistor widths (W ) in the output currentmirror, and relying on the start-up trimming to reach the requiredaccuracy. Note also how bias voltages are generated without the useof separate current branches to reduce power dissipation.

E. ECAP Amplifier

To facilitate in-vivo investigation of the physiological effects ofneural stimulation on the retina with the aim of continuously improv-ing the visual perception delivered by the implant, the implant mustbe able to measure Electrically evoked Compound Action Potentials(ECAPs) and send the measurements across the transcutaneous link.The first stage of an amplifier designed to measure such signals isshown in Fig. 8. As with ECAPs recorded in other implant systems(e.g. [14]), the magnitude is expected to be in the range 10μV to1mV sitting on a large residual voltage (the stimulation artifact;in the range 100mV to 10V). In order to implement an amplifierwith good noise performance at low power dissipation, the amplifieris implemented wholly with low-voltage transistors while the large

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Fig. 7. Trimable current DAC bit with short turn-on time. Unconventionaltransistor sizing is used to lower parasitic capacitances: transistor dimensions,W/L in μm are shown next to each transistor.

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Fig. 8. First stage of ECAP amplifier. Phase 1: reset phase where residualelectrode voltages and amplifier off-set are sampled; phase 2: amplificationphase.

stimulation artifact and electrode voltages are sampled across on-chiphigh-voltage capacitors (HV) in a reset-phase [15].

IV. FURTHERWORKIn addition to the circuits presented, a number of system design

measures are taken to reduce the overall power dissipation in theimplant. Firstly, all inactive circuits (e.g. unused ECAP amplifiers)must be turned off and their clocks gated. Secondly, electrodevoltages and VDDH level should be monitored and fed back to theexternal parts for closed-loop control over the transmitted power(i.e. VDDH should be set as low as possible). Thirdly, stimulationalgorithms that use a minimum number of stimulations to conveythe image are being developed (as an example, one could imaginesending only image outlines rather than the whole image).In addition to power dissipation control and electrode shorting,

we are also working on implementing other new, secondary safetyfeatures. One example of this is a leakage current monitor. The onlyallowed electrical connections between the tissue and the implantelectronics are via the electrodes. Faults in the implant encapsulation,e.g. due to trauma, can establish other electrical paths which maythen cause harm. We have designed a low-power leakage detector,which monitors the power supply currents to the circuits driving theelectrodes (positions marked “x” in Fig. 3) and raises an alarm ifthose are different [16].With the exception of digital control circuits and an ADC, the

components presented in this paper are the core functions that weplan to integrate in our epiretinal implant. At the time of writing, thecircuits are in various states of development. The hexagonal electrodestimulation arrangement, the inductive link and the forward datachannel have all been tested experimentally. The other componentsare currently being designed in a 0.35μm HV CMOS technology.

V. CONCLUSIONSIn this paper, we have discussed the design of vision prostheses for

restoring vision to the blind. We argued that a most promising systemapproach is to design an epiretinal intraocular neuro-stimulator who is

fed power and data from an externally worn system. We presented ahexagonal arrangement of retinal stimulating electrodes which canbe driven concurrently with little interaction between stimulationsites and whose driving electronics can be scaled to the hundredsof electrodes required for useful vision restoration. We identifiedvolume constraints and power dissipation as the most critical designconstraints in the implant. Thus, we outlined the design of a numberof core circuit functions we are currently working on which willenable the implementation of a epiretinal neuro-stimulator.

ACKNOWLEDGEMENTSThe authors wish to thank Dean Wheatley and Wu Jing Yifor valuable discussions and the Australian Research Council forfinancial support.

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