136 IEEE TRANSACTIONS ON ELECTRON DEVICES, VOL. ED-29, NO. 1 , JANUARY 1982

eters of the fluid and the sensor. However, the measurement pleted with on-chip circuitry except for the transistor which results can be corrected by making use of the information maSureS the flow t€mPeratUre. about the flow temperature which is measured by the second chip. ACKNOWLEDGMENT

The authors wish to thank E. Smit of the IC Workshop at the Delft University of Technology for the technological sup-

V. CONCLUSIONS port rendered to the flow-sensor project. A monolithic integrated flow sensor has been realized which

is suitable for measuring gas and liquid flow velocities. The sensor is direction-sensitive in one dimension, but can also be [ l ] made sensitive in two dimensions when one wishes to measure the velocity vector in a plane parallel to the chip surface. The [21 sensor has a square-root-like static response, which is dependent on several flow parameters. Measurement results have been f 3 1 given for air flow in the range from 0 to 3 m/s at room tem- [41 perature. I t can be expected that further developments will lead to a group of relatively simple and inexpensive flow trans- L 5 1 ducers. Although the present sensor consists of three transistors [61

on a chip with external circuitry, future sensors can be com-

REFERENCES L. A. Rehn et al., “Dual-element solid-state fluid^ flow-sensor,” Society of Automotive Engineers, Inc., Tech. Rep. SAW/SP-80/

G. E. Platzer and M. Southfield, “Solid-state fluid flow sensor,” US Patent 3 992 940, Nov. 23, 1976. A.F.P van Putten, and S. Middelhoek, “Integrated silicon ane- mometer,” Electron. Lett., vol. 10, pp. 425-426, 1974. R.W.M. van Riet and J. H. Huijsing, “Integrated direction-sensitive flowmeter,” Electron. Lett., vol. 12, no. 24, pp. 647-648, 1976. A. J. Chapman, Heat Transfer. New York: McMillan, 1974. R. M. Warner and N. N. Fordemwalt, “Integrated Circuits, design principles and fabrication,” in Motorola Series Solid-state Elec- tronics. New York: McGraw-Hill, 1965.

~

458/S02.50, pp. 101-106,1980.

Thin-Film Multielectrode Arrays for a Cochlear Prosthesis

SHIHAB A. SHAMMA-DONOGHUE, GERALD A. MAY, MEMBER, IEEE, NEIL E. COTTER, ROBERT L. WHITE, FELLOW, IEEE, AND F. BLAIR SIMMONS

Abstract-The design and fabrication of flexible thin-film microelec- trode arrays for use in a cochlear prosthesis are described. The elec- trode array is designed to be inserted through the round window of the cochlea into the spiral scala tympani chamber of the cochlea. A life- time of decades under stimulation is sought. The electrode array is comprised of photolithographically defined platinum-on-tantalum conductors sandwiched between polyimide layers. A liquid polyimide is used, which polymerizes in two stages. After the first stage of curing, the polyimide is susceptible to photolithographic etching, allowing patterned access holes to be cut into the top layer of the insulating sandwich. After the second cure, the polymer becomes inert biocom- patible Kapton. The processing techniques and the electrode test results are presented.

Manuscript received May 4, 1981; revised September 3, 1981. This work was supported by the National Institutes of Healthunder NINCDS Contract N01-NS-2336.

S. A. Shamma-Donoghue, N. E. Cotter, and R. L. White are with the Department of Electrical Engineering, Stanford University, Stanford, CA 94305.

T I. INTRODUCTION

HE LOSS OF HEARING is a personal tragedy best appre- ciated by the more than 200 000 people in the United

States alone suffering from profound neurosensory deafness. At present, there is nothing science or medicine can offer to restore their hearing, and it is for them that a cochlear audi- tory prosthesis is being developed. This auditory prosthesis is aimed at profound deafness caused by a nonfunctional inner ear (cochlea).

The structure and function of the cochlear prosthesis is shown in Fig. 1 referenced to a normal functioning ear. A normally functioning ear receives information in the form of air pressure variations which are collected by the outer ear and transmitted by the middle ear to produce vibrations in the fluid-filled inner ear. The vibrations are then converted by special sensor cells into electrical pulse patterns which

G. A. May was with the Department of Electrical Engineering, Stan- travel up the many fibers of the auditory nerve to the brain.

View, CA 94043. ford University, Stanford, CA. He is now with Cromemco, Inc., Mt. The prosthesis attempts to bypass a defective inner

F. B . simmons is with the ~ i v i ~ i ~ ~ of Oto~aryngo~ogy, Stanford Uni- ear by direct injection of electrical signals into the auditory versity School of Medicine, Stanford, CA 94305. nerve [ 2 ] . A microphone detects the sound patterns and

0018-9383/82/0100-0136$00.75 0 1982 IEEE

SHAMMA-DONOGHUE e t al.: MULTIELECTRODE ARRAYS FOR COCHLEAR PROSTHESIS 137

b !MPEDANCE I . F lLTERiNG . 1. TRANS- 1 lNFORMAT!ON PROCESSING MATCHING ' D iSTRlBUTlON I 'DUCTION I -

INNER EAR

A C O U S T I C REFLEX

N O ? M i L E A R

AUDITORY PROTHESiS \

!SK!N

Fig. 1. Block diagram of the auditory system and the auditory pros- thesis (after Dallos [ 11).

AUDITORY N E R V E

NERVE FIBERS

S

M O D I O L A R ELECTRODE ARRAY

w \ \ \\I\

\ 3' IGAN OF C O i i T l

(b) Fig. 2. Highly schematized drawings of the rigid modiolar array inserted

into the eighth nerve (a) and of a flexible scala tympani array inserted into the cochlea (b).

delivers the signal to a speech processor. The speech processor extracts the significant speech information from these signals and organizes the information into a coded set of electrical signals which the brain finds meaningful. The electrical sig- nals are then injected into the auditory nerve either directly, using an electrode array inserted into the auditory nerve through a hole drilled in the modiolus, or indirectly using an electrode array inserted into the lower chamber of the cochlea, the scala tympani, and stimulating the nerve endings as they approach the basilar membrane of the cochlea. These two modes of signal injection, and the type of electrodes used to accomplish them are shown in highly schematized form in Fig. 2 . This paper is concerned with the stimulating micro- electrode array which interfaces the electronics to the neural system, the final stage of the prosthetic system.

Stimulation of the auditory nerve by externally generated

electrical currents to produce sound sensations was first dem- onstrated nearly two centuries ago. The generation, however, of controlled, complicated, predetermined, and intelligible auditory sensations through direct electrical stimulation has only recently become possible, due both to advances in elec- tronics technology and to advances in the knowledge of audi- tory physiology. A major obstacle to the development of a successful auditory prosthesis has been the difficulty of fabri- cating reliable and reproducible multichannel microelectrode arrays small enough to reproduce the normal electrical activity on the cochlear nerve by stimulating distinct subsets of nerve fibers. In the case of the modiolar arrays, these difficulties have been largely overcome through the development of tantalum-on-sapphire microelectrode arrays [3] .

We have since turned our attention to the development of a scala tympani microelectrode array. This paper presents a new flexible Pt-Ta-on-polyimide array, fabricated with the high reproducibility and dimensional accuracy afforded by planar photolithography and other integrated circuit techniques.

11. G E N E R A L REQUIREMENTS O F T H E S C A L A TYMPANI ELECTRODES

The scala tympani electrode must meet quite rigorous elec- trical, mechanical, and biocompatibility requirements. Elec- trical stimulation must be delivered in a localized manner and with sufficient intensity to elicit supra- threshold percepts, but without degrading the electrode through electrolysis or damag- ing the neural tissue through excess current or charge densities. Mechanically, the array must be stiff enough to be pushed up the scala tympani without buckling, yet flexible enough to follow the curved spiral structure of the cochlea, and it must slide up the scala tympani without damaging the delicate basilar membrane separating the scala tympani from the scala media. Finally, the structure must be fully biocompatible, both in the passive sense that all materials used must be appro- priate for long-term chronic implantation without adverse tissue reaction and in the active sense that passing electrical currents through the electrode interface must not release toxic ions into the surrounding tissue. We consider these constraints sequentially in the following subsections.

A . Electrical Requirements One of the first design decisions to be made is the size and

spacing of the stimulating electrodes.. We address first the electrode size, the area of the stimulation current source. On the one hand, we would like the electrode as small as possible, since we want localized stimulation. On the other hand, reducing the electrode area increases the current or charge density which must be delivered through the electrode. Neither the current density nor the charge density can be increased arbitrarily, since they are limited by both the sus- ceptibility of metal electrodes to electrolysis at high charge delivery densities and by the susceptibility of the surrounding tissue, neural or otherwise, to damage at high charge delivery or current levels. We, therefore, must establish how much charge the electrode must deliver to achieve the required percepts, and then adjust the electrode area so that charge and current densities are kept below damaging levels.

We show in Fig. 3 a plot of perceived loudness versus charge

1 3 8 IEEE TRANSACTIONS ON ELECTRON DEVICES, VOL. ED-29, NO. 1 , JANUARY 1982

LEGEND: 0 - ELECTRODE 1, M Hz 0 - ELECTRODE 2, 50 HZ A - ELECTRODE 3. 50 HZ SUBJECT: JM

’ i o ’ 20 ’ i o ’ 40 50

Fig. 3 . Plot of perceived loudness versus charge injected for a modiolar electrode array. The subject was asked to rate loudness on a scale on which 100 cents was maximum comfortable loudness.

CHARGE ( n C l

injected for a human subject with an electrode array implanted in the modiolus.’ Note that thresholds for perception are on the order of 5-10 nC/phase of pulsatile stimuli, and that maxi- mum comfortable loudness is achieved with charge injections on the order of 50 nC/phase. For pulse durations less than 200 ps/phase, the loudness perceived is determined primarily by the total charge injected rather than by current density or voltage. We expect that charge injection required for threshold or for maximum loudness will be somewhat higher for monopolar scala tympani electrodes than for modiolar electrodes because the relevant neural tissue is more remote. Experiments in other laboratories suggest the difference can be as great as a factor of two. We, therefore, accept as a design criterion that the individual electrodes must be capable of delivering 100 nC/phase of pulsatile stimulation.

We now ask how small electrodes capable of such charge injection can be made. This size is bounded by electrochem- ical considerations on the electrode material and by neural damage considerations which are independent of the electrode character. We first must decide on the electrode materials, since that choice determines the electrochemistry involved. The first class of materials useful for biostimulation electrodes, and the class universally used in actual practice to date, is that of the metals, especially the noble metals; platinum and iridium. The second class, fairly recently arisen, is that of oxide capacitors, such as tantalum pentoxide [4], which deliver charge to the tissue by electrochemically passive charging and discharging of an inert layer. Though this second class shows great promise, the capacitative electrodes tested to date do not show sufficient charge delivery capability for our application, so we will consider only metal electrodes. We will, in fact, consider only platinum, since it has been validated by a number of investigators as one of the metals (iridium and rhodium being the others) most resistant to deterioration by electrolysis.

lThese data were generated by L. Atlas and M. Herndon in experi- ments at Stanford University with implanted subject RF (May 15, 1978).

1 .o t

C H A R G E D E L I V E R E D pC /cm2 Fig. 4. “Charging curve” of platinum-physiologic saline interface,

showing electrochemical processes involved. The curve shows that between 300 and 400 pC/cm2/phase can be delivered before irrevers- ible processes are encountered.

Electrical stimulation of. the nervous system involves the transfer of charge from a metallic electronic conductor to an ionic liquid conductor-a mixture of low molecular weight ions and large protein molecules in biological saline. For small charge transfer, the process can be purely capacitative, involving the charging and discharging of a charge double layer present at the metal-electrolyte interface. At the charge- transfer densities required for functional neural stimulation, however, one or more faradic reaction usually come into play [3] . Extensive electrochemical investigations, notably by Brummer and his associates [SI, [ 6 ] , have resulted in an aced rate characterization of the platinum-electrolyte interface. Fig. 4 shows the electrochemical “charging curve” for the platinum-electrolyte interface. The most important conclu- sions to draw from this figure is that charge densities on the order of a few hundred microcoulombs per square centimeter can be delivered through the interface without entering the region of irreversible surface reactions such as the liberation of O2 or H2 and that such charge delivery is achieved at an interface voltage of less than 2 V.2 Biphasic charge-balanced pulses tend to reduce the effects of surface electrochemistry, but perfect reversibility is achieved only if all reaction prod- ucts are immobilized at the surface and recaptured on the next half cycle.

Brummer et al. [7] have determined that a platinized and aged [8] platinum-saline interface can safely deliver up to 300 &/geometric cm2 of charge per phase at currents up to 300 mA/geometric cm2. For bright (smooth) or electro- plated platinum surfaces these values are somewhat smaller, due to smaller roughness factors (real area/geometric area). If we choose 100 pC/geometric cm2 as our maximum design charge delivery density, well within Brummer’s criteria, and assume the 100 nC/phase charge requirement for maximum comfortable perceptual loudness, then we arrive at an elec- trode area of 1 X cm2, approximately 300 pm X 300 pm. Platinum dissolution rates at these charge densities’ are very low, between 1 and 50 ng/C. An electrode of 6-pm thickness,

2These reactions are discussed somewhat more fully in [ 3 ] .

SHAMMA-DONOGHUE e t aE.: MULTIELECTRODE ARRAYS FOR COCHLEAR PROSTHESIS 139

SCALA MEMBRANE TYMPANI

. . . .

COCHLEARIS

. . . . . . . . . . . . .

(c)

Fig, 5. Drawing of the cochlea to illustrate the mechanical constraints on a scala tympani electrode. (a) Cochlea as seen from middle ear. (b) Cochlea with part of wall removed to show basilar membrane. (c) Cross section of one turn of the cochlea, showing its subdivision into compartments,

stimulated at 150 pps for 16 h a day at maximum charge den- sity should last approximately 5 years, assuming most pessi- mistic dissolution rates.

The data on maximum charge density or maximum current density for neural damage are much less complete and much less compelling. Best indications are that the neural damage criteria are similar to the electrode dissolution criteria. In o’ur own experience, chronic stimulation of our human subjects at charge delivery levels of up to 50 nC/per phase for up to 3 years has produced no evidence (such as increased thresholds) of neural damage.

B. Mechanical Requirements The mechanical requirements are difficult to quantify, but

reduce’ to the requirement that the array be practical to insert up the scala tympani, preferably through the round window of the cochlea, without trauma to the fragile membranes or linings of the scala tympani.

Some appreciation of the mechanical problems involved can be obtained by reference to Fig. 5 which presents in Fig. 5(a) a drawing of the cochlea, viewed from the direction of the middle ear, in Fig. 5(b) a drawing of the same cochlea with part of the wall removed to show the basilar membrane, and in Fig. 5(c) a cross section of one section of the cochlea, showing the transsecting membrane. The cochlea are drawn four times actual size. The scala tympani electrode is designed so that it can be inserted through the round window (88 in Fig. 5(a) or 5(b)) and one and one-half or two turns up the scala tympani, a distance of perhaps 2 cm. The scala tympani cavity is roughly 1 mm in diameter at its widest and tapers to 0.5 or 0.6 mm in diameter in the apical turns. The aspect ratio of the electrode must, therefore, be greater than 20: 1. It is a long thin electrode. Further, the electrode must not

rupture the basilar membrane separating the scala tympani from the cochlear duct or the scala vestibuli; extensive de- generation of the spiral ganglion cells, whose axons are the auditory nerve, is known to result from such trauma. The electrode array must also not damage the stria vascularis, the lining of the scala tympani, since evidence is accumulating that such damage also causes neural degeneration.

Most practical scala tympani arrays up till the present have consisted of fine wire bundles in molded silastic carriers. For such electrode arrays, the mechanical and electrical properties are necessarily interrelated. As is discussed below, meeting electrical and mechanical requirements on the scala tympani array using such technology has proved very difficult, and it is probably fair to say that a scala tympani array of truly satis- factory mechanical properties has yet to be accomplished.

C. Biocompatibility It is essential that all elements of the electrode array be bio-

compatible. The preferred materials are Teflon, silicone rub- bers, Kapton, certain epoxies (Hysol A), platinum, iridium, rhodium, tantalum, titanium, Parylene. Pyre-ML and Mylar are biocompatible, but their ability to withstand protracted saline immersion is suspect. The electrode metals must be utilized in such a fashion that they do not liberate potentially toxic ions into the cochlea.

111. HISTORICAL BACKGROUND ON SCALA TYMPANI ELECTRODES

To date, the only viable scala tympani electrode arrays have been fabricated from bundles of fine wires. The difficulties of controllable and reproducible fabrication of such electrodes are considerable, primarily because of the small dimensions involved. These electrodes are presently assembled in many different ways and with varying degrees of complexity. The techniques range from simple single-electrode wires [9] to multiple wires inserted through individual holes drilled in the cochlear walls [ 101 , to multiple fine platinum-rhenium electrode wires and flamed ball or flush-cut tips held in a molded silastic carrier [ 1 I ] .

There are two basic disadvantages associated with the fine- wire bundle technology:

1) There are coarse limits to the minimum electrode geom- etries that can be produced reliably and in a controllable manner. All but the single-wire electrode have dimensions that are large relative to the scala tympani chamber causing a displacement of the chamber’s fluids when inserted. This is suspected to have a detrimental effect upon long-term neural survival.

2) In all fine-wire bundle electrodes, the electrical and mechanical properties of the array are tightly coupled. The number of elements in an array cannot be readily extended while maintaining the same overall dimensions or flexibility.

An attractive alternative to these methods of fabrication is the use 6f photolithographic integrated circuit techniques which can readily resolve the problem of control and repro- ducibility. The required stimulation areas (300 pm X 300 pm) are well within typical photolithographic precisions. The application, however, of integrated circuit techniques to the fabrication of flexible scala tympani arrays is not straight-

140 IEEE TRANSACTTONS ON ELECTRON DEVICES, VOL. ED-29, NO. 1, JANUARY 1982

, 2 5 0 x 2 5 c u CONTACT UOLES

Fig. 6. Schematic drawing of a polymer-platinum-polymer thin-film scala tympani electrode.

forward. The materials, both conductors and insulators, commonly used in industry are inappropriate and unsatis- factory for implantable biostimulation electrodes. The metals, gold and aluminum, are subject to electrolysis and, in the case of the latter, have toxic effects. SiOz insulation softens and deteriorates when soaked in saline fluids (body fluids).

Nevertheless, attempts have been made in recent years to fabricate scala tympani arrays using integrated circuit tech- niques [12] . Typically, these electrodes were fabricated using conventional polymeric films such as Teflon or Mylar to form polymer-metal-polymer sandwiches. The technology of fabrication was not pursued to the point where the prob- lems of conductor continuity, metal adhesion, and lead attachment were solved or even substantially addressed.

In manufacturing the new array, an internally consistent fabrication sequence has been developed based on the choice of materials which, while uniquely suitable for integrated circuit fabrication, satisfy the functional requirements of biostability and long-term reliability. The electrode is an eight-element Pt-Ta array sandwiched between two thin layers of an insulating flexible polymer (polyimide), Fig. 6. The processing and the choice of the electrode materials, together with the reasons behind their assumption and devel- opment are discussed below.

IV. THE CHOICE AND PROCESSING OF THE POLYMER

In order to be used in the fabrication of a flexible electrode by means of photolithographic techniques, the “supporting polymer” must have the following properties:

biocompatibility and biostability, and chemical inertness; mechanical flexibility in order to be inserted into the spiral structure of the cochlea; insolubility in body fluids; impermeability to the ionic species of the body fluids; relatively low dielectric constant, reducing interelectrode crosstalk and capacitive leakages; freedom from pinholes and cracks; good metal adherence to support the actual metal elec- trode line under stimulation in body fluids;

8) IC processing compatibility: the polymer has to have processing properties compatible to photolithographic IC processing techniques to be utilized. These include good adherence to commercially available photoresists and precision etchability using wet or dry etchants.

Pyralin (PI-2555) [13], a polyimide (Kapton), has experi- mentally been found to satisfy most of these requirements. This polymer has the distinctive advantage of being available in preprocessed liquid form to be used in a manner paralleling that used with photoresist. It has a two step curing cycle. After the first cure cycle the polymer is solid, but is still chemically reactive and susceptible to photolithography and chemical etching. At the end of the second, it acquires all of the desirable properties chemical inertness, water insolubility, and ion impermeability.

This polymer belongs to the family of “Aromatic Polypyro- mellilimides,” derived from “Aromatic Polyamic Acids.” When in liquid form, the polymer is a soluble polyamic acid. After the first curing step, most of the solvent is driven off and the polymer is partially polymerized. Complete polymerization (cross linking) occurs only after a second high-temperature (350°C) curing step.

Pyralin PI-2555 is one of a large family of polyimides of dramatically different terminal properties after curing. For pyraline, the diamine constituent of the polyimide is an oxide, and the terminal polymer after complete curing is Kapton. Kapton has been proved biocompatible, and shows excellent thermal and hydrolitic stability under harsh conditions [ 131 . A list of the important properties of Pyralin, supplied by the manufacturer (DuPont) is given in Table I.

The viscosity of the preprocessed liquid polymer can be adjusted by adding more of the solvent; however, curing procedures have to be adjusted accordingly. An optimum cure cycle was experimentally determined with polymer layer thicknesses from 25 to 100 pm.

50°C 3 h 80°C 4 - h solvent

100°C 4 - h evaporation 135°C 1 h 300°C 1-h polymerization.

Several time-step sequences were tried for the solvent evaporation step. It was found that the best evaporation cycle is one that increases the temperature of the evapora- tion at a rate which minimizes the formation of large gas bubbles from the underlying layers of the polymer. The temperature at which the solvent is completely evaporated is 135°C. When more rapid curing cycles were used, it was found that these gas bubbles would evolve, resulting in craters and extensive pitting of the surface. The curing cycle pre- sented here results in a very smooth polymer surface. For thicker layers (greater than about 200 pm), it is advisable to extend the last bake to longer times (even overnight). It was also found that to prepare thin layers (<lo pm) suitable for photoresist patterning and alkaline etching, it was necessary to cut the entire first baking step to a single 90°C bake for

SHAMMA-DONOGHUE e t al.: MULTIELECTRODE ARRAYS FOR C ‘OCHLEAR PROSTHESIS 141

TABLE I PROPERTIES OF PYRALIN PI-2555 __

Liquid

Viscosity 10-12 poise

Density 1.06 g/cc

Solvent N-methyl-2 pyrrolidone aromatic hydrocarbon

Flash point 64OC

F i l t r a t ion 1 micron

Solid Film

Tensi le s t rength

Elongation

Density

F lex ib i l i t y

Melting point

Final decomposition temp.

Dielectric Const. ( 1 kHz)

Dissipat ion factor

Die lec t r ic s t rength

Volume r e s i s t i v i t y

19,000 psi

10%

1.39 g/cc

180’ bend, no cracks

None

560‘C

3.5

.002

4000 vol ts/mi 1

1016 ohm-cm

5-10 min and then to proceed with photoresist application as follows:

Baking, Patterning, and Etching of Thin Polymer Films ( I 0 -20 pm)

Spin polymer liquid on a substrate (such as a Si wafer). Bake at 90°C for 5-10 min until apparently dry. (Solvent

Spin positive resist on polymeric film. Bake at 60°C for 15-20 min. Develop the photoresist and simultaneously etch polymer.

Positive resist “developers” are corrosive alkaline solutions which attack the exposed polymer as well as the exposed positive resist and etch both away.

will not be completely removed at this stage.)

Postbake at 90°C for 1 h, then remove resist by acetone. Follow the bake sequence indicated earlier for thicker films.

V. Pt-Ta CONDUCTOR ARRAY In the scala tympani electrode, the metal conductor array

is contained between two layers of polymer. The “bottom” layer is thick (50-150 pm) giving the electrode an adequate balance of strength and flexibility. The “upper” layer is thin

Fig. 7. Stress patterns of thin Ta films.

(10-20 pm) and insulates the conductor array from its fluid environment while allowing stimulation and bonding at open- ings at both ends of the film (see Fig. 6). The method of sputtering and patterning the metal films is discussed in the next section.

The conductor array was conceived initially to be of Pt only; properties of this metal, described in Section I1 [3], made it attractive as a stimulation interface in the modiolar electrode. Platinum, however, is known for its poor adhesion to many substrates due to its inert (noble) qualities. Hence, in order to promote the Pt-polymer bond, it is necessary to presputter the polymer surface for at least 10 min in an argon ambient. This treatment eliminates a weak boundary layer at the poly- mer surface, a region lacking a well defined transcrystalline structure. According to Schonkorn [ 141 , cross-linking occurs tit the surface of a large number of polymers when electron- ically excited rare-gas ions are allowed to impinge on the surface [ 171 .3 This cross-linking increases markedly the co- hesive strength of the surface by the formation of a dense gel matrix. Such surfaces support much stronger adhesive bonds than the untreated polycrystalline surface.

Despite a significant improvement in Pt-polymer adhesion after argon sputter-cleaning, the adhesion is still inadequate. Pt films under continuous stimulation (25 &/phase, 100 pulses/s) lift off in less than a week (3-5 days) while others, soaked in saline without electrical stimulation, persist for many weeks.

The search for a solution led to the use of an intervening layer of Ta. Ta is a much more reactive metal and can, when freshly sputtered, be caused to bond to the highly polarized -CO bonds of the polymer. Initial results were discouraging as 3000-6000-8 films of Ta could not be made to adhere to the “bottom” polymer under various sputtering conditions. The patterns, .however, of tantalum liftoff were interesting and strongly indicative of stress in the film. A large number of “parallel sine waves,” sometimes symmetrically radiating from a point defect in the film, appear in various regions of the specimen (see Fig. 7).

Stress relief patterns like these are generally related to the

3Cross linking in polyethelene, for example, occurs as the metastable and ionic gases contact the polymer surface causing the abstraction of hydrogen atoms. The resulting polymer radicals interact to form cross links.

142 IEEE TRANSACTIONS ON ELECTRON DEVICES; VOL. ED-29, NO. 1 , JANUARY 1982

1 S I L I C O N WAFER 1 C L E A N W A F E R

/ Lioulo ' P O L Y M E R / 5 1 L I C n N W A F F R . . . - . . - . .

S I L I C O N W A F E R //,SOL,ID /POLYMER+'

S P I N I D R I P T H I C K POLYMER LOWER LAYER

F IRST CURE SEQUENCE SECOND CURE SEQUENCE

A N D P L A T I N U M SPUTTER TANTALUM

SILICON W A F E R CONDUCTOR PATTERN LITHOGRAPHY & E T C H I N G

POLYMER LAYER S P I N / D R I P U P P E R

FIRST CURE OF U P P E R P O L Y M E R

L I T H O G R A P H Y & ETCH OF UPPER POLYMER

F I N A L C U R E SIL ICON REMOVAL

Fig, 8. Schematic drawing showing process sequence for fabrication of flexible thin-film scala tympani electrode.

presence of localized tension and compression areas on films with ,moderate to weak adhesion [ 151 . Thinner Ta films (1000 A) still exhibited this phenomenon but only after pro- longed soaking in saline. Only at thicknesses of 300 a or less was the stress pattern absent, due probably to the incomplete formation and transition of the film from the f.c.c. to b.c.c. structure. In such transitory films Chopra has consistently found these particular patterns (above critical thicknesses depending on substrates and sputtering conditions).

Using 300-8 tantalum films as intervening layers between Pt and polymer has improved the adhesion substantially. This is particularly true when combined with an O 2 presputtering etch (instead of argon) for 5 min to transform and oxidize the surface of the polymer. Thinner layers of Ta may also be used, but extended experiments with Ta films of 50-200 a have indicated a definite improvement in the long-term relia- bility of the adhesive bond for the thicker films. The opti- mum film thickness is approximately 250-300 A.

VI. SUMMARY OF PROCESSING STEPS The processing sequence illustrated in Fig. 8 starts out with

a bare Si wafer, upon which a film of the liquid polymer is spun. The film is then baked to a thickness of 50-100 pm as required. A thin 300-8 Ta film is sputtered following a 5- min O2 presputtering step. The film is deposited in an RF diode sputtering system with a cathode voltage of - 1.5 kV in an argon ambient. A Pt film (6000 A) is then sputtered onto the Ta film. The conducting electrode pattern is then formed photolithographically using RF back-sputtering with the photo- resist as a protective mask. The remaining positive resist is stripped very gently with acetone and a second layer of polymer (10 pm) is spun onto the wafer and prebaked (as

* * I , &* * "

Fig. 9. Photograph of a flexible thin-film scala tympani electrode.

TABLE I1 ELECTROPLATING OF PLATINUM*

Pre-clean

1 . 20-30 second cathodic treatment in a 10-20% ac id so lu t ion ( e . s . H2S04)

2. Quick r inse in 2% H2S04.

3. P u t i n to p l a t ing so lu t ion w i th cu r ren t ON and s l i g h t a g i t a t i o n fo r 10 seconds.

P la t ing Condi t ions

1 . Use "platinum T P " p l a t i n g s o l u t i o n (Amino n i t r a t e s a l t i n Amnia) .

2 . Use a "iieavy" bath schedule:

a . 12 mi. T P s o l u t i o n

b . 26 m l . de-ionized water

c. temperature: 1 2 0 ° F (50'C)

d . p l a t ing cu r ren t dens i ty = 5 . 4 mA/r.crn2

3. P l a t i n g r a t e = 2 . 5 :rn/hr

*Plating solution is obtained from Technic, Inc., ST1 6100. PlatinumTP.

described earlier). In order to open stimulation and contact holes, a positive resist mask is formed and the polymer is etched at the required sites to expose the metal. After re- moving the positive photoresist, the array is baked at 200°C for 2 h to fuse the two layers completely. The arrays are then cut to the required shape using a 0.004-in-thick diamond saw. The polymer film normally comes off the Si wafer at this stage. If necessary, a quick dip in HF : HN03 1 : 1 etches the silicon leaving the intact array behind (note that this etch does not affect the polymer or the exposed Pt). A photo- graph of a completed electrode array is shown in Fig. 9. Finally, 2-mil, Kapton-insulated Pt wires are welded to the contact pads using a spot-welding technique. The strength of the bond is a sensitive function of the welding technique, requiring tight control over the direction, energy, and pressure of the welding pulse. The welded joint is then covered with a drop of the liquid polymer which is subsequently baked to fuse to both the electrode and wire insulation.

When under stimulation, Pt dissolves into the surrounding fluids at a very small rate (1-50 ng/C). Because of the long implant duration proposed ( 5 years), it is necessary that a thicker layer of Pt (-6 pm) be deposited at the stimulation sites. This is typically best done using well established electro- plating techniques discussed elsewhere. Due to the small areas involved, fine control is necessary over the cleaning procedure and the currents employed during the deposition. Typical cleaning and deposition schedules are shown in Table 11.

VII. TESTING THE ELECTRODES Electrode lifetime under chronic stimulation must be com-

mensurate with the time scale of the overall auditory pro- sthesis. Both in vitro and in vivo tests have been carried out;

SHAMMA-DONOGHUE et al.: MULTIELECTRODE ARRAYS FOR COCHLEAR PROSTHESIS 143

in vitro tests to monitor the integrity of the array components, looking for evidences of degradation such as insulation de- terioration, films peeling off, impedance changes, dissolution of stimulation pad platinum; in vivo tests primarily to examine biocompatibility, tissue reaction (with and without stimula- tion), and neural trauma from insertion. In Vitro Testing: The scala tympani array has been subjected

to continuous current stimulation at moderate levels while soaked in saline solution. The basic test setup is simple and allows a monitoring of injected charge and voltages developed across the electrode interfaces and solution. The state of the stimulation metal pad interfaces is inferred from a variety of indicators such as voltage waveform parameters-ohmic drop (being dependent on the electrode geometric area and hence insulation integrity) and initial interface capacitances (initial slope of potential build up reflecting the electrode real area, and hence surface condition)-and careful examination by light and electron microscopy.

As discussed earlier, initial tests exposed problems of Pt- polymer adhesion which were overcome through the use of a Ta intervening layer. These electrode arrays have success- fully passed 10 weeks of continuous testing during which no adhesion deterioration was observed. All waveform parameters remained stable, indicating no observable structural changes at the Pt-saline interface and hence confirming the visual evi- dence of electrode integrity. These results, however, cannot be readily projected for longer periods. Only prolonged life testing can determine whether the Ta-polymer bond is ade- quately strong and durable. Such tests are now in the process of being automated through the use of a microprocessor- controlled testing station.

In Vivo Testing: Histological studies are being initiated in cat preparations to test for the electrode material’s long-term biocompatibility and ’diostability under normal stimulation conditions. One of the initial problems that had to be ad- dressed in these studies was the neural trauma resulting from electrode insertion in the scala tympani. Neural trauma has at least two sources. Rupturing the delicate basilar membrane during insertion is known to cause subsequent degeneration of auditory nerve fibers. Pressure of an electrode array against the walls of the scala tympani, excluding the normal fluids, has also been shown to result in nerve degeneration, probably through local nutritional and vascular deprivation. The larger molded fine-wire bundle arrays often cause both kinds of damage. Our thin-film planar arrays appear to cause damage primarily by tearing or cutting of the basilar membrane with their sharp edges.

In order to overcome the latter problem, we have opted to modify the overall flexibility of the array and to blunt its sharp edges by the use of a silastic carrier. This carrier can be made in very small dimensions since it contains no wire bundles; i.e., the electrical mechanical properties of the array are still completely separated and no limitations are imposed by one upon the other. The mechanically modified array has been inserted up to 13 mm in cats’ scala tympani with no apparent damage inflicted, as determined by fixing and sub- sequent sectioning of the cochlea with the electrode inserted. At writing, long-term chronic life tests in cats were about to commence.

VIII. CONCLUSION This paper has described the design and fabrication of a

scala tympani microelectrode array which serves as the charge delivery interface for a cochlear prosthesis. The array is fabri-‘ cated using integrated circuit photolithographic techniques which allow significant miniaturization with a high degree of accuracy and reproducibility. The electrode consists essen- tially of a platinum conducting array sandwiched between two layers of polymer. Platinum adhesion to the polymer is pro- moted by the use of a thin intervening layer of tantalum. The conditions necessary to achieve adequate platinum-polymer adhesion are outlined in detail; in particular, the restrictions on the tantalum film and the presputtering treatment of the polymer surface. The polymer chosen for this application is the polyimide Pyralin (PI-2555). This polymer satisfies the necessary requirements of being mechanically flexible, elec- trically insulating, and uniquely compatible to integrated circuit processing.

The successful fabrication of this array offers an important alternative to the traditional wire-bundle electrodes. Initial in vitro and in vivo tests have demonstrated the reliability of the arrays with regard to metal adhesion and insulation integrity. These tests will continue over the course of the next six months before it is definitely known whether such arrays are suitable for scala tympani implants in human subjects.

~ - ~~ . . - ~ ~

ACKNOWLEDGMENT The authors are indebted to Dr. F. T. Hambrecht for his con-

tinuing support, both fiscal and intellectual, of this work.

REFERENCES P. Dallos, The Auditory Periphery. New York: Academic Press, 1973. T. Gheewala, R. Melen, and R. White, “A CMOS implantable multielectrode auditory stimulator for the deaf,” IEEE J. Solid- State Circuits, vol. SC-10, no. 6,1975. G. May, S. Shamma, and R. White, “A tantalum-on-sapphire microelectrode array,” IEEE Trans. Electron Devices, vol. ED-26, no. 12,1979. D. L. Guyton and F. T. Hambrecht, “Theory and design of capacitor electrodes for chronic stimulation,” Med. Biol. Eng., pp. 613-620, Sept. 1974. S. B. Brummer and J. McHardy, Current Problems in Electrode Development in Functional Electrical Stimulation, F. Hambrecht and J. Reswick, Eds. New York: Dekker, 1977, pp. 499-514. S. Brummer and M. J. Turner, “Electrical stimulation of the nervous system: The principle of safe charge injection with noble metal electrodes,” Biochem. Bioenergetics, vol. 2, pp. 13-25, 1975. S. Brummer, J. McHardy, and M. J. Turner, “Electrical stimula- tion with Pt electrodes: Trace analysis for dissolved Pt and other dissolved electrochemical products,” Brain Behav. Evol., vol. 14, p. 10, 1977. G. Hills et al., Reference Electrodes. New York: Academic, 1901. W. House, “Multichannel electrical stimulation in man,” in Elec- trical Stimulation o f the Acoustic Nerve in Man, M. Merzenich et el., Eds. San Francisco, CA: Velo-bind, 1974. C. Chouard et al., “Implantation of multiple introcochlea elec- trodes for rehabilitation of total deafness: Preliminary report,” The Laryngoscope,vol. 86, no. 11, 1976. E. Hochmair et al., “An eight-channel scala tympani electrode for auditory prostheses,” IEEE Trans. Biomed. Eng., vol. BME- 27, 1980. See also review in R. White, “Development of multi- channel electrodes for an auditory prosthesis,” Tech. Rep. in

144 IEEE TRANSACTIONS ON ELECTRON DEVICES, VOL. ED-29, NO. 1, JANUARY 1982

response to RFP-NIH-NINCDS-80-05,1980. p. 1373. See also, P. Shah, “High performance, high density [12] P. Clark and R. Hallworth, “A multiple electrode array for a MOS processing using polyimid interlevel insulation,” in Proc.

cochlea implant,” J. Laryngology and Otology, vol. XC, no, I , 1976.

IEEE 23rd Electric Components Conf. [I41 H. Schonhorn, “Surface modification of polymers for adhesive

M. Sonn and W. Feist, “A prototype flexible microelectrode bonding,” in Polymer Surfaces, by D. Clark and W. Feast, Eds. array for implant-Prosthesis applications,” Med. Biol. Eng., New York: Wiley, 1978. p. 778, Nov. 1974. [15] K. Chopra, Thin Film Phenomena. New York: McGraw-Hill,

[ 131 Developed by Dupont in 1965. See J. Polymer Sci., pt. A, vol. 3, 1969, p. 311.

Thin Linear Thermometer Arrays for Use in Localized Cancer Hyperthermia

Absfract-A thin linear array of silicon diodes has been developed to measure one-dimensional temperature profiles in tissue during treat- ment of cancer by localized hyperthermia. The array is composed of six discrete diodes on three flexible stainless steel wires, with a maximum cross-sectional dimension of 0.5 mm, so that it can be introduced into a tumor area via a small puncture wound. Temperature data are ex- tracted using an external electronics system under microprocessor con- trol; the present overall accuracy of the system is 0.2OC over the range of 37-45’C. The array has been tested in ultrasound, RF, and micro- wave heating fields. Computer simulation shows this array to be non- perturbing of the thermal field in tissue, so that it can provide accurate temperature data. Development of a batch-fabricated array of twenty diodes on five leads is under way. A monolithic silicon diode array is shown to produce large temperature perturbations because of its high thermal conductivity, while an alternative technology using silicon micromachining and adaptations of existing techniques for lead fabri- cation should produce an array of low thermal conductivity which can obtain accurate measurements. The present and future arrays should also be suitable for data collection in many in-vivo situations other than hyperthermia.

L I. INTRODUCTION

OCALIZED HEATING (hyperthermia) of cancerous tissue has been shown by several investigators to produce shrink-

age of tumors when used as the sole treatment modality and when used in conjunction with either chemotherapy or ionizing radiation [ l] - [ 6 ] . Hyperthermia is accomplished using ultra-

Manuscript received June 3, 1981; revised September 3, 1981. This work is part of the research of the National Resource for Silicon Bio- medical Transducers, sponsored by the NIH Division of Research Re- sources under Grant NIH-RR-01086.

The authors are with the Department of Electrical Engineering, Stan- ford University, Stanford, CA 94305.

sound, RF, or microwaves for deep heating of tissues, often in conjunction with heating or cooling of the skin to produce a peak temperature at a given depth below the surface of the body. One of the basic problems in hyperthermia is determin- ing the amount and extent of temperature rise in tissue. Great variations in tissue structure, vascularization, and blood flow occur from individual to individual, at different body locations within the same individual, and at the same body location in an individual over the course of time; as a result, identical treatment regimens with a given heating apparatus can produce wildly different temperature profiles and unrepeatable results.

Successful use of hyperthermia in a clinical setting will re- quire feedback in the form of temperature measurement, at several points within the tumor and in the tissue surrounding the tumor. Ideally, such temperature measurement should be noninvasive, so that no disturbance of tissue occurs due to the temperature measurement. In practice, some minimal amount of tissue invasion is necessary if temperature is t o be measured at several localized interior points. At a minimum, a single puncture wound is necessary. If several “thermometer points” can be located along the length of such a puncture, a practical optimum will be achieved.

This paper discusses the development of a miniature linear thermometer array composed of silicon diodes on flexible wires which will provide a near-optimal means of temperature measurement during hyperthermia. This array should also prove suitable for use in many other in-vivo situations. For example, temperature profiles through the heart wall and on the surface of the heart in a normal experimental animal are of interest to cardiologists, and should be obtainable using arrays of this type.

0018-9383/82/0100-0144$00.75 0 1982 IEEE