600

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 39, NO. 6 , JUNE 1992

Computerized Transcutaneous Control of a Multichannel Implantable Urinary Prosthesis Mohamad Sawan, Member, IEEE, FranGois Duval, Magdy M . Hassouna, Jin-sheng Li, Mostafa M . Elhilali, Joel Lachance, Member, IEEE, Marc Leclair, Soheyl Pourmehdi, and Jaouhar Mouine

Abstract-In the present paper we describe a personal computer interface of a multichannel implantable urinary prosthesis. This system is composed of two main parts: the first one is internal and consists of an implant using a 4-pm CMOS gate array chip controlling a wide variety of waveforms via eight monopolar channels. The second, an external controller featuring a versatile software, a PCB card plugged in a portable microcomputer, and a radiofrequency-coupled technique. This device is used to transmit the power, the data and the synchronization clock to the implant by a simple binary signal modulating a 20 MHz carrier. We also report the features of implant encapsulation and electrode design and fabrication. In the experimental phase, we studied the effect of early electric stimulation of the bladder during the spinal shock phase in the dog. We present the operative techniques that enabled us to perform chronic electrostimulation of the sacral roots and discuss the results.

INTRODUCTION T the beginning of the 20th century, the bladder sensitivity to the electrical stimulation had been discovered [l]. Until 1971, however, the research pursued did not report any promising results. In fact, the majority of the investigated techniques [2]-[4] consisted in stimulating the bladder muscle (the detrusor). The published results were not favorable because this kind of stimulation involved a nouniform contraction [5] and required an extensive electric current spread that entailed other complications; pain, outlet blockage, tissue injuries, and destruction of electrode tips caused by fibrosis [6]. Many researchers proceeded to stimulate both detrusor and sphincter [5], with even less interesting results since this modality implied an external sphincter closure during bladder contraction. An important step was done by Burghelle [7] and Einsenberg [8] who started to stimulate

A

Manuscript received December 6 , 1989; revised November 12, 1991. This work was supported by the Kidney Foundation of Canada, the Natural Sciences and Engineering Research Council of Canada, the Fonds de la Recherche en Sant6 du Ouebec. M. Sawan is with the Department of Electrical Engineering, Ecole Polytechnique, Montreal, P.Q., Canada H3C 3A7. F. Duval, J . Lachance, M. Leclair, S . Pourmehdi, and J. Moui'ne are with the Department of Electrical Engineering, Universite de Sherbrooke, Sherbrooke,. P.Q., Canada. M. M. Hassouna, J . Li, and M. M. Elhilali are with the Department of Urology, McGill University, Montreal, P.Q. Canada H3A 1 A l . IEEE Log Number 9108126.

the bladder through the nervous pathways [7]-[9]. This modality of stimulation showed that less energy is required to command the bladder (voiding or retention). Walter [ 101 proposed the volume-conduction method which is effected by the conduction of electricity through the fluid medium. Many research groups [ 111-[ 151 have been reporting considerable progress; the most important is the localization of the micturation controling nerves and the access to those nerves through the skin or by some surgical techniques. The exact mechanism of the bladder control, however, is still not clearly identified. A special difficulty is the separation of the sacral nerves functions. On the technical side, the majority of commercially available stimulators lack some of the following necessary options: 1) independent multichannel implant, 2) wide range of programmable parameters, 3) high efficiency in energy and data transmission, 4)friendly user's interface, 5 ) high-frequency stimulus generation, 6) waveform flexibility (monophasic, biphasic, etc. . . .). Table I shows the most important systems. The majority of them are based on the AM transmitter. The demodulator extracts the data from the received RF signal and sends it directly to the tissues via the electrodes. Gorgis's stimulator is an adapted pacemaker. Its most important characteristics is the magnet control used to turn it on and off. The Medtronic stimulator is programmable by means of two apparatuses and contains a telemetry circuit to confirm the data validation. The complexity of the human machine and the weakness of the available stimulators incited researchers to design new multichannel transcutaneous and fully programmable implants [21]-[23]. These devices are still under development. The available techniques and materials either to encapsulate implants or to make biocompatible electrodes present important deficiencies, especially the lack of a flexible interface to tissues. The radiofrequency-coupled devices still present some restrictions-that is, a limited energy -. transkission capability-which require the implant localization inside the body to be near the skin. Despite the numerous discoveries, the progress Until today has not been satisfactory in either area (electrostimulation and surgical). The need of a versatile electrostimulation system to study urology dysfunctions is urgent. In

0018-9294/92$03.00 0 1992 IEEE

__

SAWAN er a l . : TRANSCUTANEOUS CONTROL OF A URINARY PROSTHESIS

60 1

TABLE I MOSTWIDELY COMMERCIALIZED TRANSCUTANEOUS STIMULATORS Characteristics

Capacity

Stimulus

Item

Investigator*

P

F

L

C

E

Particularity

References

Two types. Max. frequency 200 Ha. Programmable, battery magnet control. Programmable, magnet control, telemetry. 3 inductive coupling Max. frequency 300 Hz. 8 parallel outputs (same information).

1151%1161

1

Avery@

M

I

9 V

1

2

Gorgis@

M

I

8mA

1

1 2 1

3

Medtronic@

M

I

10.5 V

2

2

4

Finetech@

M

T

25V

3

3

5

Physico-Med@

M

I

Unk

1

8

P (Polarity) F (Format) L (Maximum level)

C (Number of Channels) E (Number of Electrodes) M (Monopolar)

lnductlve COUPllnS

-Cantroller

Sldn

Implant

Flg. 1. Block diagram of the complete electrical stimulation system

this paper we propose a complete experimental system (Fig. 1) which is composed of an IBM PC or compatible portable microcomputer, a microcomputer hardware interface, a data interface stage for inductive coupling, a voltage converter and regulator, and finally a gate array integrated circuit. The system has the following advantages: capability to prepare and generate a wide variety of stimulus, good efficiency and high displacement tolerance of the radiofrequency-coupled inductive link (24% at 25 mm distance between transmitter and receiver antennas), an 8-monopolar channel fully programmable neural prosthesis, new materials to encapsulate implants, and a new technique to produce biocompatible electrodes. It is possible that this system, with only minor modifications, could be used for other neurostimulation applications (cochlear stimulation, pain control, electrotactile stimulation, etc. . . . ). The modifications consist in software algorithm adaptations and adjustments of some passive elements on the hybrid circuit to set the maximum current and the transmission data baud rate. The system constitutes a useful tool to investigate many tests in the neurourological domain. In addition, the outside part is the preliminary version of a miniaturized controller. MICROCOMPUTER SOFTWARE INTERFACING The software part of the system described here consists of a multifunction program that allows the user to com-

1171 1181

1141. 1191

WI

I (Impulse) T (Train) Unk (unknown)

municate the necessary information to the stimulator. All the input-output tasks (data control and transfer, keyboard decoding and display, etc. . . . ) are programmed in assembler language, whereas data analysis and processing tasks are developed in Pascal. The desired stimulation parameters can be either directly communicated from the keyboard or read from an existing file. In fact, the interactive style of this software, which allows a great flexibility to operate the system, is characterized by four distinct option menus (Table 11). The first menu (MENUl), appearing automatically on the screen, comprises six options dealing with stimulations. The option “STIMUL” leads to the third menu (MENU3) which allows us to specify the stimulating strategy before processing and dispatching the data. The second menu (MENU2) encloses six options and can be reached by the first-menu option called “MENU2”. This one allows us to manipulate files that enclose stimulation parameters. The stimulation parameters of interest are: the current amplitude, the pulse (high frequency) with its duty cycle, and the train (low frequency) with its duty cycle (Fig. 2). They can be introduced by using the option “PARAM” of the first menu. The fourth menu (MENU4) reached by the “TEST” option is dealing with three test options mainly developed to operate new circuits. The first one (“TESTl”) provides data for any one of the current levels entered by the user (level: 0 to 63) following any stimulation strategy. The second option (“TEST2”) allows the user to test command words by sending them as separate bytes. The last option (“TEST3”) allows the generation of a sawtooth current waveform and performs bipolar stimulation. The maximum current level of stimulation is limited at 2 mA. In fact, special attention has been paid to control the active phase of the stimulating current. This current is checked every time that new parameters are entered to satisfy a simple relation that defines the maximum charge quantity [23], [25]. If the current is wrong, the user is informed and has a chance to change the parameters. Ta-

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 39, NO. 6, JUNE 1992

602

TABLE I1 DIFFERENT MENUS OF THE MICROCOMPUTER SOFTWARE INTERFACING Items

Option

MENU 1

Description

ISTOP

Halt the program and retum to DOS system. Begin stimulating. Initialize the system. Enter or modify the stimulation parameters (table 11). Access to tests menu (MENU 4). Access to files manipulation (MENU 2). Halt the program and retum to DOS system. Save the entered parameters. Get parameters from an existing file. Display existing files directory. Erase an existing file. Return to first menu (MENU 1). Generate stimuli on both pairs of electrodes. Generate stimuli by alternating the pairs of electrodes. Generate stimuli on only one pair of electrodes. Retum to the first menu (MENU 1). Generate data with current level specification. Enter a 24-bit command word and then generate it once. Generate bipolar stimulation using a sloped current wave.

2 STIMUL 3 INIT 4 PARAM 5 TEST 6 MENU 2 1 STOP 2 STORE 3 RETR 4 DIR 5 ERASE 6 MENU 1 1 SIMULT 2 ALTERN 3 ONlOFF 4 MENU 1 ITEST 1 2 TEST 2 3 TEST 3

MENU2

MENU 3

MENU4

.

. l a w frequency

Hlgh and duty f m wcyck ncy

and duty cycle

Fig. 2. Description of stimulation parameters: stimulus waveform

TABLE 111 THESTIMULATION PARAMETERS’ EXTREME LIMITSOF ONE CHANNEL PAIRS Stimulus Frequencies

(Hz)

Duty Cycles (%)

Amplitude (PA)

High

Low

High

Low

Pulse Width (ms)

Q/Phase (PC)

0-2000

1-4500

1-1000

10-90

10-90

0-900

0-900

Q: Charges quantity of the active phase

ble 111 depicts the stimulation parameters’ extreme limits variation of each bipolar channel. Note that the charges quantity ( Q ) is the area resulting on the exponentially current variation. It can be approximated by 50% of the product of the current amplitude and the pulse width.

MICROCOMPUTER HARDWARE INTERFACING Precise circuit details for this printed circuit board are not given, since the actual circuitry involved uses standard digital CMOS integrated circuits and logic configurations to reduce power consumption. This device will be operated with any IBM-PC (XT or AT) or compatible computer. All timing is derived from the microcomputer internal clock which, when divided synchronously, provides both

Data bus

=T 12 I ul&

PB

Control bus

%

Address bus

I !1

clwd

4.77 MHz

Fig. 3 . Simplified block diagram of the microcomputer hardware interfac ing.

a bit clock at the transmission rate (300 kHz) and a 24-bit word clock (9 kHz). As shown in Fig. 3 , the present card receives three data bytes and loads them into two PPI (Programmable Peripheral Interface). The 9 kHz word

SAWAN

1'1 U / . :

TRANSCUTANEOUS CONTROL OF A U R I N A R Y PROSTHESIS

clock controls the transfer of each command word to the parallel to serial shift register (24 bits), which is clocked serially into the Manchester encoder by the 300 kHz clock. The Manchester encoder integrates the data and the synchronization in a simple digital signal [22], [24]. This coded data is transmitted to the implant via the data interface stage.

603

LLu I 1-77

I

Power

regulator

Coded Data

skin

U

Modulated carrier

F

Fig. 4. Block diagram of the data transmission stage.

DATAINTERFACE STAGE In this section we deal summarily with the power and data link. We note that on the block diagram of the whole system, the external and internal electronics are coupled through an inductive link. Although there are many approaches for sending information to the implanted stimulator, we have retained this one because of its greater safety, efficiency, and high tolerance to displacement, compared with systems using optical detectors or acoustical approaches [26], [27]. In addition, infection is avoided since the skin is not broken, as is the case with the percutaneous link. Thus, the transmitter designed for sending information leaves the skin intact and transmits power indefinitely at field levels low enough to meet the government safety standards (10 mw/cm' [28]). Fig. 4 depicts a simplified block diagram of the data transmission stage. It encloses a 20-MHz AM modulator that produces a carrier the amplitude of which is modulated by the incoming coded data (Manchester code). The resulting modulated carrier is then presented to a class D amplifier, the output of which is controlled by a power regulator. This device is intended to compensate for the variations of the coupling factor effected by the displacements between the transmitting and receiving coils. It increases the power voltage of the class D amplifier in the case of high coupling and it reduces the power voltage in the reverse case [Fig. S(a)]. The secondary output signal of the transformer is then presented to a voltage regulator and an AM demodulator to provide coded data and power to the whole implant. With such a system, two main objectives are sought: a minimum transmission bandwidth of about 1 MHz, since it has been designed for a maximum transmission rate of SO0 kbits/s; and minimum losses to ensure the extemal power source durability of the future miniaturized system. To reach these two objectives, we have opted for a stagger-tuned link. Together with the power regulator technique, this link features implied low energy consumption [Fig. S(a)] and possesses the advantage of a wide bandwidth [29], [30] in addition to its adaptability to coils displacements. It also produces a fairly stable output voltage from which to recuperate a regulated voltage supply. The voltage regulator used in our system operates with a small voltage drop across it, so that it consumes little power while sourcing large currents, thus satisfying the second objective mentioned above. In the same optic, let us note that the choice of the 20 MHz carrier depended on the transformer efficiency since a computer simulation model

Charge impedance variations (01

(a)

0

5

10

15

20

25

30

35

Axial distance (mm)

(b) Fig. 5. Radio-frequency coupling results: (a) power regulator, (b) power transmission efficiency.

and experimental tests [Fig. S(b)] have indicated a maximum efficiency (24%) at this frequency [30].

MULTICHANNEL CHIPAND STIMULATION STAGE The main part of the implant described here consists of a powerful microprocessor chip developed to perform versatile control of different desired stimulation algorithms. Intended for many neuromuscular stimulation applications, it has been adapted to bladder stimulation. Table I11 shows the main characteristics of the chip, but their main features are described in other papers [24], [2S]. Moreover, we detail its different aspects that enabled us to realize a bladder controller. This device, controlling eight monopolar (or four bipolar) channels, consists of a low-power CMOS 4-pm gate array integrated circuit. It works as a microprocessor executing 24-bit command words at a transmission rate of 300 kbits per second. At this transmission rate, a total control of the stimulation waveforms (shape, frequency, amplitude) is effected by means of a relatively large number of command words. Other characteristics are shown in Table IV. Fig. 6 shows the essential parts of the chip. The Manchester decoder extracts the data and clock from the received digital signal. The synchronization block verifies the 3-bit header and then loads the data into the 21-bit data register. The command decoder synchronizes the data flow among the control units, which command the output

IEEE TR.ANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 39, NO. 6, JUNE 1992

604

TABLE IV MAINCHARACTERISTICS OF THE MULTICHANNEL CHIP Item

Parameter

Comment

Technology Die dimension Gates number Operation frequency Command word Channels number Stimulation speed Power voltage Maximum consumption

I S 0 CMOS 4-Micron Gate Array 6x6" 2000 1 MHz 24 bits 8 monopolar (4 bipolar) 0-4.5 kHz 3.0-7.0 V 8 mw L Fig. 7. Hybrid circuit of the implantable prosthesis.

able circular hybrid circuit (Fig. 7) which is then installed in our new urological stimulator when sealed for a longterm implantation (more than five years) in a circular titanium box.

Coded slgnal

I GND

Fig. 6. Simplified block diagram of the CMOS gate array chip.

stage of the chip. This output stage is an analog circuit composed of a 6-bit digital to analog converter: the current level can take any one of 64 possible discrete values. In manufacturing, an external resistor is trimmed to set the upper level of the current scale. The current sources intensity for a charge impedance of 1 kQ, at a maximum current level of 2 mA, have shown linear variation characteristics. In addition, the maximum current level of 2 mA remains stable for a wide range of charge impedance variations (100-2500 Q ) . In order to reduce the electrolysis process occurring at the surface of the electrodes and thus protect the biological tissues, balanced stimulus should be used to neutralize the charge distribution around the electrodes [31]. In addition to coupling capacitors to balance the flow of the charges, the control unit of each current source commands some commutators to generate biphasic waveforms. To bring the chip into operation, note the following necessary pins and functions:

RC:

This pin allows the extraction of the clock from the coded digital signal; so it is intended to connect the necessary resistor-capacitor network to the Manchester decoder designed inside the chip. POL: This pin holds the necessary reference voltage to set the maximum current intensity. EO.,: These pins are the output of the eight channels of the system. H / B " : This pin selects one of two chip devices. This configuration allows us to get 16 channels. The chip and necessary circuits are mounted on a suit-

STIMULATION ENVIRONMENT AND ELECTRODES DESIGN Electrodes are intended for a hostile body environment, so they have to be corrosion resistant and unconditionally biocompatible. Tantalum, as a conducting metal, is a particularly interesting choice [32]-[35]. In vitro and in vivo tests have demonstrated that tantalum produces less effects on tissue than do silver and stainless steels [32]. In addition, it is easier to insulate it electrically from saline fluid than gold and platinum [32], [33]. In fact, a thin layer of tantalum pentoxyde (Ta205), eminently biocompatible, can be grown on tantalum surfaces by an electrolytic process [36]. To avoid any adverse effects on tissue, however, the electrode-nerve interface must be assured by noble metals (platinum, iridium, palladium, etc . ) for their inert characteristics [34], [35]. Gold contacts processes should be avoided for long-term applications because of their low resistance to electrolysis; however, platinum contacts are known to offer, like iridium and rhodium, excellent resistance to electrolysis under conditions of balanced stimulus. On the other hand, in order to attenuate tissue damage and hydrolysis, a large electrode-nerve contact, reducing the current density, is necessary. Moreover, applying biphasic stimuli can optimize reversible reactions. Finally, to avoid the friction between the electrode and the body environment, a thin layer of medical silicon [37] for final insulation of the electrode will reduce the imtation and decrease the current leakages [35]. The noble metal surface is usually unsealed, but a very thin membrane should be applied to avoid irritation risks. Electrodes Design Fig. 8 shows the different steps that we used to make electrodes. Tantalum foils are cleaned, an! a solution of citric acid is used to help grow a 1000 A of tantalum pentoxyde (Ta205). At one end the electrode tip (1 X 2 mm) is prepared to make the electrode-nerve contact. On

SAWAN et al.: TRANSCUTANEOUS CONTROL OF A URINARY PROSTHESIS

-

-

\

-

10 MII (a)

Gold connector Gold tube

605

/ d

*

12 cm

(b) 1.6"

t'

(5"

t

43"

I

t

(C) 0 SUastic UllTantalum IGold ( 1 micron) 0 Ta205 [io00 A )

I

Fig. 8 . Electrodes design: (a) Tantalum foil ready to make extremities contacts. (b) Mechanical connector and tube fixed on the electrode tip. (c) Completed sample monopolar electrode.

5 cm

-

Fig. 10. Mechanical assemblage and encapsulation of the whole implant.

its radio-frequency permeability, required by the use of an RF link between the implanted device and the external processing unit [38]. A low-cost approach satisfying these criteria has been developed in our laboratory, using polymers as encapsulating material. The encapsulation technique shown in Fig. 10 consists of six principal steps:

--

I-

I

U

0

Silastic

IGold

I1 micron)

(d)

Fig. 9 . Electrodes categories: (a) monopolar type, (b) bipolar with Silastic foil as interface, (c) bipolar with Silastic tube as interface, (d) bipolar with two parallel outputs with Silastic foil.

the other end of the electrode (1 X 10 mm), a gold-plated connector is mechanically fixed, together with a goldplated tube to avoid breaking of the tantalum electrode at the level of the junction. The final step of the process consists of encapsulating the electrode for biological insulation. This step is done by surrounding the electrode with Silastic tubing and moulding. With this process many types of electrodes were fabricated. Fig. 9 illustrates one monopolar, two bipolar, and one double bipolar electrode configurations. This last one is used to stimulate two separate nerves using the same stimulus. Finally, three electrode-nerve contact types have been inve2tigated: gold sputtering (0.001 mm), gold plating (4000 A), and platinum welding-a foil of platinum welded by electrical discharge on the foil of tantalum. This technique provides a good mechanical and electrical configuration of electrode-nerve area. Our investigations are now turned to this material. PACKAGING Encapsulation of the implanted device involves three goals: biocompatibility , mechanical rigidity, and electrical insulation of the circuit. Its biocompatibility will avoid the rejection of the implant by the human body, while the mechanical rigidity and the electrical insulation will assure the good functioning of the device for long-term implantation. Another important quality of the packaging is

Drop a thin layer of wire-bound protector on the hybrid circuit. Fix the receiving antenna on the thick hybrid. This is a circular one-inch diameter antenna consisting of three rounds of copper wire (or one round of platinum wire). Thread flexible stainless wires into a biocompatible rubber tubing. Then crimp the wire to a gold-plated connector and mould an insulating support around this connector. This is made of a medical-grade elastomer (Silastic). Stainless wires, when recovered by Silastic, provide good biocompatibility [39], which results in long-term protection from the body environment. In addition, we used this wire for its flexibility and workability. Place the wired connectors to crimp, before hand welded, to the output pad of the hybrid plate. Mould the whole hybrid in a low-shrinkage epoxy. We chose the Araldite 502 resin with 951 hardener [40] because of its biocombatibility, low water absorption, high hardness after polymerization, and because it offers a good mechanical support for the antenna and wires. This epoxy permits the embedding of wires outputs, which is needed to avoid water infiltration between the Silastic tubing and Araldite casing. Finally, as we did with electrodes, we coat the prosthesis with Silastic. This last precaution provides the maximum biocompatibility of the implanted device. Notice that the transparency of those three polymers allows for easy detection of encapsulation defects and direct observation of the final product (Fig. l l ) . Practically, polymers have the major disadvantage of absorbing water after several months. In order to guar-

606

lEEE TRANSACTlONS ON BIOMEDICAL ENGINEERING, VOL. 39. NO. 6, JUNE 1992

-

2 Bipolar channel

Stimulator

----

Sphincter

PeMc floor

Fig. 12. Implant and electrodes disposition. Fig. 1 I . Photograph of a complete implant

d antee proper functioning of the device for long-term implantation, we expect to encapsulate the implant into a titanium case; this metal has already been successfully tested for medical applications with heart stimulators [41]. EXPERIMENTAL METHODA N D RESULTS In the McIntyre Animal Center at McGill University, ten dogs weighing between 10 and 25 kg have received implants. In order to study the effect of early electrical stimulation of the bladder during the spinal shock phase on the return of detrusor activity, the operative approach comprises the following steps [15]: the animal is turned into the prone position, and two laminectomies are done in the same sitting: dorsal and sacral. The dorsal laminectomy is meant to expose and cut the spinal cord at the level of D9-10. The sacral laminectomy exposes the cauda equina with the upper sacral roots. By meticulous dissection, in the spinal canal, each sacral root on each side is identified and individually stimulated with a bipolar nerve probe connected either to a Grass signal generator (MODEL 79D) or to a locally built stimulator. The parameters of stimulation are selected. The changes in the bladder and urethral pressures are observed during the stimulation of each root. The roots that give the highest intravesical with sometimes the lowest intraurethral pressure are identified. The locally available tantalum electrodes were wrapped around one root and connected to the implant (Fig. 12). Using two bipolarchannel implants, many configurations were tried; similar electrodes were wrapped around two roots, one electrode was wrapped around one root and another connected to the subcutaneous tissues near the implant as a return electrode. The implants are located in a subcutaneous pouch on the flank of the animal. The electrical stimulation was carried out at least twice a day even during weekends and holidays. Cystometrograms (CMG), using medium-rate filling with normal saline, were done weekly for the first month post operatively. The animals had three postoperative work-ups

h e preoperative one. They were kept for a period of four to eight months. Some of the animals were not given implants; they were used as a control group, with voiding by means of intermittent catheterization. The clinical parameters used for comparisons between the two groups of dogs were the following: serum creatinine, BUN (Blood Urea Nitrogen), IVP (Intravenous Pyelography), VCUG (Voiding CystoUrethrography), and CMG (Cystometrogram). The parameters of stimulation are the pulse amplitude (0.5-2.0 mA); the train waves are composed of a high-frequency signal of 1-1600 Hz modulated by a low-frequency signal of 1-500 Hz. The stimulation technique has been applied either intermittently or continuously during 10-40 s.

Results The stimulation was done from the first postoperative day, and the residual poststimulation volume was maintained at a very low level ( < 20% of the capacity). The maximum intravesical pressure ranged between 50.and 70 cm H 2 0 . There was movement of tail and hind limbs, but the procedure was well tolerated by the animal. After the first few weeks, the stimulation was carried out only twice a day since the residual amounts were negligible. The CMG was done every week for the first postoperative month, then once a month. The radiological investigations showed normal upper urinary tract and absence of vesico-ureteral reflux. On CMG, the bladder capacity at voiding was between 150250 mL (mean 175 mL), with detrusor contraction during voiding. The experiment was terminated if the animal showed major complications, such as uncontrolled urinary infection and decubitus ulcers, hence avoiding any undue suffering by the animals. In all animals the upper urinary tract was intact on histological examination. In order to protect the stimulating tissues fully, we limited the quantity of maximum charges. This limitation is obtained by coupling capacitors that allow the stimulus to be free dc voltage and to generate completely balanced waveforms. Each waveform period is composed of an active phase and a charge-recovery phase. The charge quan-

__

607

S A W A N er u l . : T R A N S C U T A N E O U S CONTROL OF A URINARY PROSTHESIS

TABLE V MAINRESULTS OF THE In vivo EXPERIMENTS Configuration Item

Dog

Implant

1 2 3 4 5 6 7 8 9 10

9845 9844 9950 9889 9946 183 135 82 1 1008 1307

10 I 8 7 8 10 10

6* 4* 6*

BC 1 s1 s2 s2 s2 s2 S2 S2 S2 s2 S2

Events Unrolling

BC 2

s2 ( L & R) S2 ( R ) ha (L) Ina (L) s 3 ( R ) s 2 4- s 3 ( L ) ( L & R) h a s 2 + s3 ( L ) (R) ( L & R ) Ina (L&R) h a ( L & R ) Ina

(L&R) (L)

+

R: Right side L: Left side Start: Date of implantation Electrodes: mechanical breaking *No coupling capacitor

Start

Stop

Reasons

02- 1 1-88 01-02-89 15-02-89 08-03-89 17-04-89 03-05-89 26-07-89 30-05-90 14-08-90 1 1-02-9 1

09-1 1-88 1 1-02-89 19-02-89 09-04-89 10-06-89 12-07-89 27-07-89 04-12-90 20-04-91 05-1 1-9 1

Electrodes Electrodes Dog dead Impedance Impedance Electrodes Impedance End of period End of period End of period

S: Sacral nerve BC: Bipolar channel Stop: End of' stimulation Impedance: temporary stopping h a : Inapplicable

tity of the active phase is maintained at a value lower than 900 pC (Table 111). Resulting in vivo events are as follows: in three animals, technical malfunction (electrodes breaking) necessitated their removal after a period of some days to some weeks. In four other animals, the stimulation activities suddenly stopped after a period varying from a few weeks to several months, but normal stimulation activities were re-established one month to three months later. That allowed us to confirm that the postoperative phase implied an increase in the impedance of the stimulated area. The next two animals received implants with high-value coupling capacitors. The latest implant did not possess a coupling capacitor. Table V depicts the history of the main in vivo experiments, including the date and site of each implantation. The complete processing period is also presented. The electrode-tissue interface proved a good contact during the implantation period. Electrode breaking is due to a mechanical problem-male-connector failure of tantalum foil-which required the addition of a gold tube to eliminate this weakness (Fig. 8). Six dogs completed the postoperative period for a mean of six months. Notice that better voiding is obtained by a high-frequency signal ranging between 20 and 200 Hz, modulated by a low-frequency signal of 1-2 Hz (Fig. 13). In addition, the two-frequency combination allowed US to leave the transmitter in a fixed position during the stimulation period.

CONCLUSION We have described a new urological computerized implantable electrical stimulation system (Fig. 14). It has been very effective in bladder control. This system was designed and fabricated to be flexible, so as to be able to run a wide range of stimulation algorithms and perform many tests in vitro and in vivo in animals. This functional stimulation system has been tested over the last four years in the urology research laboratory at Royal Victoria Hospital and McGill University. It is not restricted to urol-

P.

9

HlghFrequency=ZS HZ

+ HiqhFrequency=ZOO

0 HighFrequency=1600

3 IO

e

0

40

20

60

80

100

120

Low frequency (H.1

20

.P

\

----K,

0

in ..

Q.

I o w f r e a u e n c v - I HZ

o

LowFrequenCy=SO

20

I

30

Charges QuMtitY/phw UC)

Fig. 13. In vivo electrical stimulation results in animal (dog).

Fig. 14. Photograph of the different parts of the whole stimulation system.

ogical applications, however; sufficient channels are available to perform other neuromuscular stimulation. The versatility of this stimulation system allowed us to obtain many favorable results: definition of the best voiding pa-

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 39, NO 6. JUNE 1992

608

rameters, impedance variations of the stimulating biological area, new algorithms to reduce the sphincter resistance. The reaction of the electrodes will be published soon. On the medical side, the stimulation algorithm resulted in the early return of the detrusor activity [2 11. The miniaturized version of this stimulation system is under way.

REFERENCES B. S. Nashold, H. Friedman, and R. Boyarsky, “Electrical activation of micturation by spinal cord stimulation,” J . Surg. R e s . , vol. 1 1 , no. 3, pp. 144-147, 1971. W. E. Bradley, L. E. Wittmens, and S. N. Chou, “An experimental study of the treatment of the neurogenic bladder,” J . U r d . , vol. 90, no. 5 , pp. 575-581, 1963. A. Kantrowitz and M. Schamaun. “Paraplegic dogs: Urinary bladder evacuation with direct electric stimulation,” Science, vol. 139, pp. 115-1 16, 1963. C. C. Stenberg, H. W. Bumette, and R. C. Bunts, “Electrical stimulation of human neurogenic bladders: Experience with 4 patients,” J . U r o l . , vol. 97, pp. 79-84, 1967. W. E. Bradley, G . W. Timon, and S . N. Chou, “A decade of experience with electronic stimulation of the micturation reflex,’’ Urologia Internationalis. vol. 26, pp. 283-303, 1971. H . E. Carstensen, P. S. Freed, D. A. Molony. and A. Kantrowitz. “External sphincter fatigue as an adjunct to electrical detrusor stimulation.” Investigat. U r d . , vol. 7, no. 5 , pp. 387-397, 1970. T . Burghele, V. Ichi,m and M . Demetrescu, “L’Clectroexcitation de la vessie midulaire. Etude expererimentale,” J . d’uroiogie medicale, vol. 64, pp. 317, 1958. L. Eisenberg, A . Mauro, W. L. Glenn, and J . H. Hageman, “Radio frequency stimulation: A research & clinical tool,” Science, vol. 147, pp. 578-582, 1965. W. L. Glenn, J . H. Hageman, A. Mauro. L. Eisenberg, S . Flanigan. and M. Harvard, “Electrical stimulation of excitable tissue by radio frequency transmission,” Ann. S u r g . , vol. 160, pp. 338-350, 1964. I . S . Walter, “Neural control of the urinary bladder,” Ph. D. dissertation, Univ. Health Sciences, The Chicago Medical School, 152 p., 1982. J . W. Thuroff, R. A. Schmidt, M. A. Bazeed, and E. A. Tanagho, “Chronic stimulation of the sacral roots in dogs,” European U r n / , , vol. 9. DD. 102-108. 1983. [I21 R. A. Schmidt, “Neural prostheses and bladder control,” /EEB Eng. Med. B i d . M a g . , vol. 2, no. 2, pp. 31-36, 1983. 1131 E. A. Tanagho and R. A. Schmidt, “Electrical stimulation in the clinical management of the neurogenic bladder,” J . Urol., vol. 140, pp. 1331-1339, December 1988. [I41 G . S. Brindley, “An implant to empty the bladder or close the urethra,” J . Neurol., vol. 40, pp. 358-369, 1977. A. El Rifaei, M. Hassouna, A: Fouda, R. Latt, M. Sawan, F. Duval, and M. M. Elhilali “The effect of early bladder stimulation of the spinal shock: A prelimary report,” J . U r o l . , vol. 141, pp. 10101013, Apr. 1989. -, The Micturation Stimulator. Farmingdale, NY: Avery@Laboratories 1978, 3 pp. J . C. Brocklehurst, “Noncatheter devices for urinary incontinence in the elderly,” Med. Insrrument., vol. 16. no. 3, pp. 167-168, 1982. -, Pulse Generator Magnet. Minneapolis, MN: Neuro Division Medtronic 1984. 31 pp. T. A. Perkins, “Versatile three-channel stimulation controller for restoration of bladder function in paraplegia,” J . Biomed. E n g . , vol. 8, pp. 268-271, 1986. P. Magasi and Z. Simon, “Electrical stimulation of the bladder and gravidity,” Urol. I n r . , no. 41, pp. 241-245, 1986. L. J . Seligman, “Physiological stimulations: From electric fish to programmable implants,’’ IEEE Trans. Biom. Eng., vol. BME-29, pp. 270-284, Apr. 1982. M. Sawan, M. Hassouna, F. Duval, M. M. Elhilali, J . Mouine, S . Pourmehdi, J. Lachance, M. Genest, and M. Leclair, “A new prosthetic device for fully bladder control,” IEEE-EMBS’89, Seattle, WA, Nov. 1989. M. Sawan, S . Pourmehdi, J . Mouine, Lachance, M . Leclair, M. Genest, F. Duval, M. Hassouna, and M. M. Elhilali, “Contrble tran-

scutane par ordinateur d’un implant urologique multicanaux.” CCGEI 89, Montrial, Sept. 1989. B. Nadeu-Dostie, P. Gauthier, F. Duval, and R. Abtahi, “ A multichannel neural stirnulator on a single chip,” Inrernat. Svmp. ArtiJicial Organs, Biom. Eng. Trans., Salt Lake City, UT, 1986. F. Duval, M. Sawan, J. Mouine, S . Pourmehdi, B. Nadeau-Dostie, and P. Gauthier, “ A fully programmable multichannel cochlear prosthesis, part I: Implant,” to be published. F. Duval and R. L. White, “Auditory prosthetic system for sensory deafness,” The 3 / s t ACEMB Conf., Atlanta, C A , Oct. 1978. J . D . Meindl, “Biomedical implantable microelectronics,” Science, vol. 210, pp. 263-267, 1980. E. J . Lerner. “RF Radiation: biological effects,” Special Rep., IEEE Spectrum, pp. 51-59, Dec. 1980. D. C. Galbraith. “An implantable multichannel neural stimulator,” Ph. D. thesis, Stanford University, Stanford, CA, 1985, 165 pp. P. Gauthier, “Etude et diveloppement d’un lien de transmission transcutani pour une prothCse auditive cochliaire,” M. Sc. Thesis, Universite de Sherbrooke, 1987, 137 pp. I . T. Mortimer, D. Kaufman, and U. Roessmann. “Intramuscular electrical stimulation: Tissue damage,’’ Ann. Biomed. Eng., vol. 8, pp. 235-244, 1980. J . T . McFadden, “Tissue reactions to standard neurosurgical metallic implants,” J . Neurosurg., vol. 36, pp. 598-603, May 1972. C . R. Legindy, M. Salcman, and N. Brennan. “A multiple floating microelectrode for chronic implantation and long-term single unit recording in the cat,” Electroencephalog. clin. Neurophysiol., pp. 285288, 1984. R . L. White, L. A. Robert, N. E. Cotter, and 0:H. Kwon, “Thinfilm electrodes fabrication technics,” Ann. NYAcud. S c i . , 1983. G . A. May, S. A . Shamma, and R. L. White, “A tantalum-on-sapphire microelectrode array,” IEEE Trans. Electron D o . , vol. ED2 6 , p p . 1932-1939,Dec. 1979. W. Westwood, Tantalum Thin Film. London: Academic, 1975. -, “Silastic MD-4-4210 Medical Grade Elastomer,” Dow Coming Corp., Midland, MI 1986. L. Bowman and J . D. Meindl, “The packagaging of implantable integrated sensors,” IEEE Trans. Biom. Eng., vol. BME-33, pp. 248255, Feb. 1986. J . B. Park, Biomaterials an Inrroduction. New York: Plenum, 1979. - “A Liquid Epoxy Resin: Araldite 502,” Ciba-Geigy, New York: Ardsley, 1982. D. F. Williams, Biocompatibility of Clinical Implants Materials. Boca Raton, FL: CRC Press, 1981.

.

Mohamad Sawan (S’88-M’90) received the B . S . degree in electrical engineering from the Universite Laval, P.Q., Canada, in 1984, the M.S. and Ph.D. degrees from the UniversitC de Sherbrooke, Sherbrooke, P.Q., Canada, in 1986 and 1990, respectively, and post-graduate training from the McGill University, Montreal, P.Q., Canada, in 1991. He is currently an Assistant Professor at the Ecole Polytechnique, Universitt! de MontrCal Montreal, P.Q., Canada. His research interests deal with the neuromuscular stimulators, the VLSI computer aided design, the analog and digital circuits design and the digital and analog signal processing applications.

Franqois Duval received the B.Sc. degree in engineering physics; the B.Eng. degree in electrical engineering; the M.Sc. degree in biomedical engineering; and the D.Ing. degree in neuronics. He is presently Professor of Electrical Engineering at the Universite d e Sherbrooke, Sherbrooke, P.Q., Canada. He developed a strong interest in the development of implantable neural stimulators while pursuing postdoctoral studies, then assuming the task of interim director of the Stanford Institute for Electronics in Medicine, Stanford, CA. Since then he has been very actively involved in designing and testing various neural stimulators, such as 8- and 16-channel cochlear implants, multielectrode bladder controllers and chronic pain controllers.

SAWAN (’I 01

TRANSCUTANEOUS CONTROL OF A URINARY PROSTHESIS

Magdy M. Hassouna received the M D degree from Alexandria University in 1976 and the Ph D. degree from McGill University, Montreal, P Q , Canada in 1985 Presently he is an Associate Professor, department of Surgery (Urology) at McGill University His main research and clinical interest5 are neurostimulation of the lower urinary tract and male sexual dysfunction Dr Hasaouna is a member of the American Urological Associdtion, Urodynamic Society and the Canadian Urological Association Jin-sheng Li received the M.D. degree from the Faculty of Medicine from Tianjin Medical University, Tianjin, China, in 1977, the M.Sc. degree from the Department of SurgeryiUrology from Tianjin Medical University, Tianjin, China in 1987. He will finish Ph.D. study in Department of Surgery, Division of Experimental Surgery, McGill University, Montreal, P.Q., Canada in 1992. His main interest is the effects of electrical stimulation on the sacral nerve5 in neurogenic bladder, Urodynamic study and prolonged electrical stimulation for evacuation after spinal cord injury. Mostafa M. Elhilali received the M.B., Ch.B. degree in 1959, the Diploma of Surgery in 1962, the Diploma of Urology 1963, and the M.Ch. degree in Urology in 1964 from Cairo University, and the Ph.D. degree in experimental surgery from McGill University, Montreal, P.Q., in 1969. He is presently Professor and Chairman of Urology, at McGill University and Urologist-inChief at the Royal Victoria and Montreal General Hospitals. His main research interests are in the areas of neurogenic bladder dysfunction, prostatic carcinoma, and male infertility. He is a Fellow of the Royal College of Surgeons of Canada, The Corporation of Physicians and Surgeons of Quebec and the American College of Surgeons. Joel Lachance (S’88-M’90) was born in Montreal, P.Q. Canada, in 1965. He received the B.S. degree in 1988 and the M.S. degree in 1991, both in electrical engineering from Universite de Sherbrooke, Sherbrooke, P.Q., Canada. His thesis research involved a design of a low power microcontroller for a cardiac stimulator. He joined the Centre Hospitalier Universitaire de Sherbrooke in 1991 as an analog circuit designer to develop X-ray sensors. His interest areas include microelectronics, medical electronics and rehabilitation engineeiping.

609

Mr. Lachance is a member of IEEE EMBS, IEEE LEOS, and Material Research Society.

Marc Leclair was born in the Quebec, Canada, in 1964. He received the B.S. degree from the Department of Physics, Universite de Sherbrooke, Sherbrooke, P.Q., in April 1988 and the M.S. degree in electrical engineering. His research project, with the Groupe de recherche en appareillage medical de Sherbrooke, focuses on the energy and data electromagnetic link between the transmitter and the implant of a cochlear prosthesis. His areas of interest also include laser processing for biomedical packaging, plasma spraying technology and the biocompatibility of materials for implant encapsulations

Soheyl Pourmehdi was born in Teheran, Iran. He received the M.EEA degree in electrical engineering from the Universite Paul Sabatier, Toulouse, France in 1986 and the M.S. degree in electrical engineering from the Universite de Sherbrooke, Sherbrooke, P.Q., Canada in 1989. He has done his research project with the Groupe de recherche en appareillage medical de Sherbrooke (GRAMS) and the Centre de recherche sur les communications de Sherbrooke (CRCS). He is presently at work on his Ph.D. degree in biomedical engineering and has developed a DSP-microbased for a multichannel cochlear implant. His areas of interest include speech analysis and digital signal processing.

Jaouhar MouYne was born in Tunisia in 1965. He received the B.S. degree in electrical engineering from the Universiti du Quebec a Trois-Rivieres in December 1985 and the M.S. degree in electrical engineering from the Universite de Sherbrooke in 1988. He has done his research project with the Centre de recherche 5ur les communications de Sherbrooke (CRCS). Since 1988 he has been working with the Groupe de Recherche en appareillage medical de Sherbrooke (GRAMS), and he is now applying for the Ph.D. degree in biomedical engineering. His research work focus on the hearing impairments prosthesis and more particularly the implants.