integrated sensor-telemetry system for in vivo glucose monitoring

E L S E V I E R Sensors and Actuators B 31 (1996) 147-153

Integrated sensor-telemetry system for in vivo glucose monitoring

Joseph Black a, Michael Wilkins ~, Plamen Atanasov b, Ebtisam Wilkins b,, a Department of Electrical Engineering and Computer Engineering. Unive,sily o] N,~w Mexico. Albuquerque. NM 87131. USA

h Department of Chemical and Nuclear Engineering. University ~fNew Mexico. Albuquerque. NM 87131. USA

Abstract

This paper reports the development of an electronics package designed for use with an implantable amperometric glucose biosensor. The package includes a potentiostat to bias the sensor and a telemetry unit to transmit information from the sensor to an external receiver and data- collection computer. The device components and package have been evaluate6 in vitro. Initial tests are performed with non-encapsulated electronics and the sensor partly submerged in a glucose solution. To simulate the implantation conditions, tests have been conducted with the electronics encapsulated in plastic (epoxy resin or silicone elastomer matrix) and submerged in a 37 °C water bath. These tests verify that the telemetry system is transmitting through an aqueous solution with a water mass equivalent to a medium sized dog. Calibration curves of the received signal versus glucose concentration in model solutions have been obtained. The correlation between data obtained using a traditional potentiostat and the data obtained via telemetry allows the commencement of in rive experimentation.

Keywordv: Telemetry systems; lmplantable glucose biosensors

1. Introduction

Diabetes mellitus is a disorder of the endocrine system manifested as an inability to control the blood-glucose level. A pressing need in the area of improvement of therapies for diabetic patients is the development of an implantable glucose sensor. Treatment would involve monitoring the glucose level and maintaining it within physiologically acceptable bounds through a tightly controlled diet or insulin injections. Insulin would be administered either by injection or by open- loop infusion. It is assumed (and has been demonstrated) that more tightly controlled blood-glucose levels will mini- mize the risk of long-term complications [ 1,2]. Current methods of monitoring glucose levels use urinalysis or self monitoring of blood glucose (SMBG), making use of glucose oxidase strips and reflectance meters. This method involves frequetit finger-prick blood sampling. This type of sampling, because it is unpleasant, suffers from a lack of patient compliance. In addition, it involves a risk of infection and, depending on the training of the individual performing the test, it is subject to inaccuracy, so its efficacy is in question [2]. It has been suggested that for SMBG to be effective it must be performed more frequently than twice a day [3]. In actual practice once a week is most common and the vast

Corrc,~ponding author. Phone: (505) 277 5906 (office); (507) 277 2928 (lab). Fax: (505) 277 5433. E-m.~_i!: [email protected].

0925-4005/961515.00 © 1996 Elsevi~.r Science S.A. All rights reserved SSD! 0925-4005 ( 95 ) 01805-6

majority of diabetic patients perform the tests less than once a day [4].

An implantable sensor could provide continuous or on- demand readings of the glucose level. This information, gath- ered on a wearable miniature computer, could be used to characterize the nature of glucose fluctuations in each patient and then develop a customized treatment plan to 'tightly control' the blood-glucose level.

A better management system would involve a closed-loop insulin infusion system such as the artificial beta cell. An implantable glucose sensor would be used to close the loop for a feedback syste~ based on an implantable insulin micro- dosing device (an insulin pump) to create an artificial pan- creas. An implmttable insulin pump has already been developed [5-7]. However, there are several unresolved problems with the development of long-term implantable glucose sensors.

Telemetry has been usc~J in biomedical research for almost three decades [8]. However, the use of telemetry with glucose sensors has only just recently been justifiable due to improved long-term operation of the sensors [9-11 ]. Tele- metry allows freedom of motion, which minimizes alterations in the biological functions bei ig monitored due to emotional and physiological stresses caused by methods of measure- ment that often require the subject to be attached to an immo- bile or heavy piece of equipment. If the telemetry module is also implantable, as opposed to being worn, then problems

148 J. B/art: e' e!. / Seny,,~¢ ond Actuators B 31 (1096) 147-153

associated with transcutaneous systems can be avoided. These problems include tissue necrosis and infection, anchor- ing, and maintaining viable site access. Depending on the range of an implantable telemetry system and the amount of data required, a patient would be able to lead a more normal life using an implantable system than with any of the currently available means of monitoring blood glucose.

This communication presents one approach to the electronics developed to allow the implantation of an amperometric glucose biosensor and telemetry unit under the abdominal skin of a dog for in vivo evaluation of system performance. Other electronic techniques are under evaluation and laboratory development.

Fig. 2. Circuit schematic of the integrated sensor-telemetry system: sub- circuit A, voltage source; sub-circuit B, the transmitter.

2. System d e s c r i p t i o n

The integrated system consists of an amperometric glucose sensor, potentiostat and radio transmitter. Fig. I shows the schematic of the system assembly. The potentiostat ( I ) provides a constant bias for the sensor, a three-electrode current- source device (2). The potentiostat also converts the sensor-generated current to a voltage and this voltage is then converted to frequency using the voltage-controlled oscillator (VCO) of a phase-locked loop (3). The frequency output modulates the carder frequency of an FIVl transmitter (4), and the signal is transmitted over the air via an antenna (5). Lithium thyonyl chloride batteries (Tadiron ~, size AA) are used as a power source (6). An FlVl receiver supplies the demodulated signal to a computer, which monitors the signal for stability and records the data. The circuit schematic of the integrated sensor-telemetry system is presented in Fig. 2.

2.1. Amperome t r i c sensor

The glucose biosensor is a three-terminal amperometric device. It employs enzyme-catalyzed oxidation of glucose and is based on the hydrogen peroxide measuring principle. The glucose biosensors consist of two parts: an amperometric electrode system and an enzyme micro-bioreactor. A three-

s

2

1 2 3 4 5 cm

Fig. l. Schematic view of the integrated sensor-telemetry system: I, potentiostat; 2, glucose biosensor; 3, voltage-to-frequency conve~er; 4, FlVl transmitter; 5. antenna; 6, batteries.

electrode amperometric scheme is used: a platinum wire (diameter 0.25 mm, length 4 mm) as a working electrode, a silver/silver chloride reference electrode (In Vivo Metric, Healdsburg, CA) and another platinum wire as a counter electrode. The working electrode is polarized at +0.6 V for hydrogen peroxide oxidation. This three-electrode amperometric system is directly inserted in the sensor housing, a plastic tube, face-side closed by a glucose diffusion membrane. The sensor housing is filled with the immobilized enzyme, thus serving as a micro-bioreactor. The glucose diffusion membrane is used to limit and control the flux of the substrate entering the micro-bioreactor. Two capillary plastic tubes, the inlet recharge tube and exhaust discharge tube, are used for replacing spent enzyme from the micro-bioreactor, without sensor disassembly. Refilling of the sensor is achieved using two septa via these tubes: one for injecting a fresh enzyme suspension, another for exhausting the spent enzyme.

The enzyme glucose oxidase (GOD) was immobilized on ultra-low-temperature isotropic carbon powder (Carbomed- ics Inc.) using the carbodiimide technique following a previously described procedure [ 12]. Prior to sensor preparation, the enzyme-modified carbon powder ( 100 mg) was dispersed in 1 ml of 0.1 M phosphate buffer solution (pH 7.4), containing 0.125 M KCI, fresh GOD (20 mg ml - i ) , and bovine serum albumin (42.5 mg ml- i ) using an ultra- sonic bath. In the preparation of the enzyme gel matrix, 2% glutaraldehyde was used for crosslinking the enzyme-modified carbon with bovine serum albumin.

Polycarbonate membranes (Poretics Co.) were used as glucose diffusion membranes, being pre-treated in Nation ® (Aldrich Chem. Co.) solution to diminish the influence of interference from plasma chemicals. The samples of membranes (diameter 37 mm) were immersed in Nation ® solution in a Petri dish for 1 h. They were then held vertically and dried in air for I h.

The biosensor was prepared by first filling a plastic tube, the electrode body (outer diameter 6 ram, inner diameter 4 mm, face-side covered by the polycarbonate membrane), with the enzyme-carbon suspension before gelling occurs. The three-electrode assembly is then introduced into the sen-

J. Black et al. /Sensors and Actuators B 31 (1996) 147-153 149

sor body and sealed. This biosensor was then stored overnight at 4 *C to complete the gelling process [ 12-14].

2.2. Potentiostat and telemetry

The potentiostat provides the bias voltage necessary for the sensor and converts the current prodnced in the sensor to a usable output signal. The bias voltage must be relatively constant over the range of variations expected in the operating environment. An integrated-cireuit (IC) current source (LM334) is employed to produce a bias of 0.6 V (see Fig. 2). The glucose sensor produces a current that is proportional to the glucose concentration in the range 1 nA-10 p.A. The current is converted to a voltage by an operational amplifier with high input impedance and a bias current much lower than the current being measured. The CMOS operational amplifier OAt, ICL7642 (Harris), has high input impedance, low bias current ( < 1 nA) and low power consumption ( < 30/~W). Noise and interference problems are minimized by using short conductor runs and shielding. The potentiostat circuit is separated from the transmitter and antenna by a copper ground plane on the back of the PC board.

The integrated micropower phase-locked-loop IC (CIM046) is used as a voltage-to-frequency converter. It converts the voltage signal to a square wave whose frequency is proportional to the voltage; this frequency is then applied to the FM radio transmitter.

In some of the constructed units a commercially available transmitter was used instead of the simple LC r.f. oscillator shown in Fig. 2. A Ramsey model FM-5 wireless microphone was chosen on the basis of the size, cost, and power consumption.

Various arrangements of long wire and coil antennae were considered for the transmitting antenna. A coil antenna 5 cm long and 0.6 cm in diameter incorporated in the unit body was selected because of space considerations, and to avoid direct contact of the antenna with the aqueous environment, with consequent heavy loading and instability of the transmitter.

2.3. Encapsulation qf the integrated sensor-telemetry system

Two approaches to system encapsulation were utilized: embedding in epoxy cement and incorporation in a silastic matrix. The device was embedded in a medical-grade epoxy cement with glass micro-bead filling, designed for the construction of implants (courtesy of Sandia National Labora- tories, and Gulton Co., Albuquerque, NM). The dimensions of the cemented device are 5.0 cm × 7.5 cm × 1.5 cm. The epoxy cement covers all the elements of the device, except the glucose diffusion membrane of the biosensor.

Incorporation in silastic (Silgard 186-silicone elastomer, Dow Coming) was performed in a mold and the elastomer was cured at 50 *C for two days. The size and shape of the system are the same as for that embedded in epoxy (the same

mold is used). Incorporation in silastic matrix, however, allows easier replacement of system parts (especially the glucose biosensor ). The cured silastic matrix combines dura- bility and elasticity. Portions of it can be cut out with a blade and then filled with new silastic material, which after curing adheres to the previously formed matrix.

2.4. Receiver and data acquisition

A communication receiver (ICOM model IC-R7000) was used to detect and demodulate the FM signal. Since the tests were to be carried out mainly inside a dog cage, an experiment with a large box used to model the cage was conducted with a long wire antenna wrapped in a spiral around the box.

The audio output from the receiver is connected to a shaper circuit and counter board in a 386 computer (PROLOG, BusBox ). The shaper circuit converts the demodulated signal to a clean square wave by an'lplification, limiting, and clamp- ing, and the computer then measures the frequency of the signal. Data are recorded by the computer in 20 s perkxLs. If the frequency change during every collection period is less than 5%, the average frequency for this period is written to a file along with the absolute time and the relative time since the start of the program.

The experimental setup for in vitro tests included a thermostatted water bath, where the measuring cell filled with model physiological fluid is placed. The distance between the measuring cell and the receiving antenna was from 1 to 5 m and both are placed in one room. In some experiments a conventional potentiostat (CV-IB, Bioanalytical Systems Inc.) was used as a standard and comparison for the circuit evaluation.

3. Results grid discussion

Several integrated sensor-telemetry systems were assem- bled, and tested at different stages of their assembly. Tests were performed with every unit before its encapsulation with a sensor model (dummy resistors) and with a glucose biosensor. Some units were encapsulated with a sensor model and tested to compare their frequency signal with a potentiostat calibration. Encapsulated units with a glucose sensor were tested partially immersed (only the sensor end) and totally immersed (the whole unit) in thermostatted 37 °C glucose phosphate buffer solution approximating the implantation conditions.

3.1. Calibration of the telemetry system

Fig. 3 presents calibration curves of the FM transmitter obtained with three integrated sensor-telemetry systems. Each unit is calibrated using a resistor model of the sensor. The model consisted of a variable resistor between the working electrode (anode) connection point (Fig. 2, point W) and the reference electrode connection (Fig. 2, point R), and

150 J. Black et al. /Sensors and Actuators B 31 (1996) 147-153

VOLTAGE SIGNAL AT TP 1, V 1 2 3

7 . o 1 ~ , I ' I ~" ~.o~ ~ 5.0

uJ • frequency • ~gn~ at TP Z

3.0 itransmitted

2.0 5 10 15 ZO

CURRENT THROUGH THE SENSOR, uA

Fig. 3. Calibration curves of ti=c FM transmitter obtained with three integrated sensor-telemetry systems directly as a frequency signal after the voltage-to-frequency converter (at Test Point 2) and as demodulated radio-transmitted signal.

a 10 kI l resistor between the reference and the counter electrode ( Fig. 2, point C) connections of the potentiostat. During the calibration procedure for each value of resistance applied between the reference and working electrodes, the potentio- star voltage output at Test Point I (Fig. 2, TP1 ) is measured using a 3.5 digit multimeter (BK Precision, model 2832) and recorded. The current passing through the sensor is calculated from the recorded voltage at TP1 and the value of the variable resistor.

The frequency output of the frequency-to-voltage converter at Test Point 2 (Fig. 2, TP2) is measured using a frequency counter (Startek model ATH-50) and also recorded. This frequency signal is plotted as a function of the sensor c~.lrrent (Fig. 3) for the three investigated units. A linear correlation between the frequency signal at TP2 and I:he current through the sensor model (voltage output of the potentiostat at TPI ) up to ! 2 p.A is observed for all the units (R2 = 0.999, n = 15).

The demodulated signal from the transmitter was applied to the computer data-acquisition system and its frequency output was also recorded. This radio-transmitted frequency signal for each unit is also presented on Fig. 3 as a function of the sensor model current. Deviations between the frequency measured by the frequency counter directly at TP2 ~nd that recorded by the computer never exceeded 0.4% in the linear range of the potentiostat (up to 12 p.A current). The slopes of all the six calibration curves coincide within an acceptable 0.5% error. Some difference in the intercepts on the frequency axis (up to 0.2 kHz) can be explained by differences in the hand-made inductor L (Fig. 2, sub-circuit B). Thus, the frequency/current calibration curve for every individual unit in this research was obtained and used in further calibrations.

The results presented on Fig. 3 demonstrate the ability of tbe potentiostat-transmitter system to provide a linear rela- tion between the frequency signal (directly or remotely recorded) and the potentiostat output. Encapsulation of the

units in epoxy cement or in silastic elastomer leads to some shift in the frequency output (shift in the intercept on the frequency signal axis), which was in all cases less than 0.1 kHz. This is probably due to some change in the value of the inductor L when the space within its spiral coils is filled with the embedding material, and changes in stray capacitance.

3.2. Calibration o f the sensor response to g lucose

In order to compare the transmitted data with the actual amperometric signal of the glucose biosensor, parallel measurements with conventional electrochemical equipment were carried out. In this case potentiostats and X - T recorders were used for data processing. The measurements were performed in a thermostatted cell with stirred electrolyte. Fig. 4 A presents a typical protocol of a sensor calibration test obtained as a current response to consecutive additions of glucose by using a standard potentiostat. The data are obtained with a freshly prepared glucose amperometric sensor [12-14] before its incorporation into the integrated system. It can be seen from the recording that the sensor current increases with the glucose concentration step change and reaches a stable steady-state value.

f 40 mg/dL glucose concentration step-than

2.5

2.0

~ 1,5

~ 1.Q P~

0.5

o.o ~ ~ ~ TIME OF CONTINUOUS OPERATION, hours

3,3 , ~ =; 3., ~ ~ ~ 3 o o o ~

~ 3.0

~ 2.9

2.8 ; ~ ; TIME OF CONTINUOUS OPERATION, hours

Fig. 4. Protocol of a sensor calibration test obtained as a current response to consecutive additions of glucose by using a standard potentiostat before incorporation of the sensor in the integrated system (A) and protocol of the sensor response to glucose concentration changes as a frequency signal transmitted by the integrated sensor-telemetry system (B).

J. Black et aL / Sensors and Actuators B 31 (1.o96) 147-153 151

The response time of the sensor (time to reach 95% of the steady-state signal value) does not depend on the glucose concentration, but on the glucose concentration step change. A response time of approximately 5 min is estimated when the concentration step-change is 40 mg dl - ~ (Fig. 4 A).

Fig. 4 B presents a protocol of the frequency response (via telemetry) to glucose concentration change~ for the sensor from Fig. 4 A after its incorporation in the integrated sensor- telemetry system (unit 001 ). In this case the encapsulated (in epoxy cement) unit was totally submerged in the thermostatted measuring cell. The glucose concentration is changed by replacing the phosphate buffer solution in the measuring ceil. The concentration step-change in this case is 100 mg dl - i. Both increasing and decreasing glucose concentration step changes are performed.

The remotely acquired output with the sensor in the blank phosphate buffer solution is approximately 2.9 kHz, which corresponds to a sensor current of 0.125/zA (by using the frequency/current signal calibration curve from Fig. 3). It can be seen from the protocol that the glucose concentration step-change of 100 mg d i - ' results in a frequency signal change of approximately 1.03 kHz, which corresponds to a sensor current signal change of about 0.70/zA.

The response time of the sensor does not depend on the direction of the glucose concentration changes (increasing or decreasing) and is approximately 25 min regardless of the glucose concentration. It can be seen from Fig. 4 B that the sensor signal values obtained when the glucose concentration increases coincide with those obtained with the decreasing glucose concentration.

The data from Fig. 4 are plotted as calibration curves of the sensor response versus glucose concentration in Fig. 5. The frequency signal scale is fitted to the current signal scale by using the calibration curve in Fig. 3 (unit 001, telemetry signal). It can be seen that the calibration curve of the sensor response to glucose obtained with a standard potentiostat before sensor incorporation (A) closely agrees with the calibration curve of the integrated sensor-telemetry system (B)

l o 5 15 20 mM 3.0 ' I ' ~' /~ 3.4

z.s A 3.3

~ 2.0 ~ d

3.2 g 1.5

3.1 z ~,.o g 0.5 3.0

0.01 100 200 300 400 GLUCOSE CONCENTRATION, mg/dl.

Fig. 5. Calibration curves of the sensor response to glucose obtained with a standard potentiostat before sensor incorporation (A) and frequency/glucose concentration calibration curve of the integrated sensor-telemetry system (B). Data from Fig. 4.

recorded remotely. Both calibration curves have the same slope (sensitivity) of 0.725 p.A per 100 mg dl- ~. Tbe difference in the current (frequency) axis intercept of = 0.1/zA could be attributed to the shift of the transmitted frequency signal after encapsulation of the integrated system in plastic. From the practical point of view t ~ . calibration curve obtained remotely can be used for a glucose concentration assay with an acceptable accuracy of 5%.

3.3. Continuous operation o f the integrated system

There are several factors involved in the lifetime charac- teristics of the integrated sensor-telemetry system. The operational lifetime of the biosensor is one of the most critical factors. Our earlier work demonstrates the ability oftbe sensor to monitor glucose concentration continuously for up to three months at 37 °C [ 14 ]. The lifetime of the independently functioning system depends on the battery lifetime. The transmitting circuit in the system is the principal energy-consum- ing component.

Tests were conducted to determine the battery lifetime during transmitter operation. The Ramsey transmitter is designed to operate at 1.5 V. The battery tests were performed using alkaline AA cells (Eveready ® Energizer). Separate batteries were used to operate the transmitter, to prevent loading problems from affecting the bias voltage for the potentiostat. Previous experiments showed that current within the milliampere range drawn from the battery would decrease the bias voltage. Tests results were consistent with the expected transmitter power dissipation of approximately 6.5 mW and the published capacity of the alkaline cells of 2.1 A h.

Lithium thyonyl chloride batteries (3.5 V) were used as a power source for the encapsulated units. These batteries were selected instead of alkaline batteries due to their higher capac- ;~ty and their suitability for medical-grade implants. In the case when a Ramsey transmitter was used in the system, a voltage regulator circuit was added to reduce the 3.5 V ~om the lithium battery to 1.5 V. The transmitter would be able to operate at 3.5 V, but the extra power dissipation in the transmitter would far exceed the power dissipated in the voltage regulator circuit.

Hermeticity of the encapsulated device is another major factor in performance. Special care was taken during the casting process to ensure protection of the integrated system. The sensor (particularly the sensor membrane) was kept wet to avoid possible damage and enzyme inactivation. The electronic parts of the assembly were protected from moisture. These two requirements arc in conflict, especially when embedding in epoxy resin is used. To solve this problem the encapsulation is performed in two steps: first, casting of the electronic parts of the system without a sensor connected; second, after curing placing the sensor in a space previously left for it. This two-step process is easier when silastic embedding is used because of the better adhesion of fresh silicone to already cured silastic. The hermeticity of the encapsulated

152 J. Black et al. /Sensors andActuators B 31 (1996) 147-153

Table 1 Integrated sensor-telemetry system response during seven days of in vitro continuous monitoring of glucose concentration of 100 mg dl- I

Time of Telemetred Glucose concentration continuous frequency obtained using the operation signal calibration curve (h) (kHz) (Fig. 5B) (mg dl- ) )

! 3.013 94 8 3.035 II1

25 3.002 86 33 3.028 106 49 3.011 93 56 2.997 80 73 2.994 76 80 3.052 126 97 3.014 95

104 3.003 87 121 3.001 85 128 3.017 97 145 3.027 105 153 3.029 107 169 3.014 95

3.2

3.1

2.9 ~ , . . . . . . . )

i i i ~ i i Z . 8 3 9 4 0 4 1 4 2 4 3 44 45 4 6

TIME OF CONTINUOUS Ot~RATION, hours

Fig. 6. Protocol of the integrated sensor-telemetry system response to glucose concentration increase (to 200 mg dl -I) during the second day of continuous monitoring of glucose ( i 00 mg dl- 1).

integrated system was evaluated together with parameters of the sensor response during the water bath tests.

Encapsulated integrated sensor-telemetry systems were tested to monitor glucose concentration for seven days at 37 °(2 in an agitated buffer solution. During this period the unit was totally submerged in the measuring cell fluid. The glucose concentration in the cell was 100 mg d l - J to approximate the physiological glucose level.

Table 1 presents a representative extraction of the seven- day test protocol. The transmitted frequency from the unit is listed together with the concentration of glucose calculated using the calibration curve (Fig. 5 B). .~t the time of collection of the data presented in Table 1. the glucose concentration in the measuring cell was always 100 mg dl - t.

Several times a day the solution in the cell was changed to 200 mg d l - i glucose, so that the sensor is alternately immersed in 100 and 200 mg d l - t glucose solutions as a test for the signal reproducibility. Fig. 6 presents a portion of a

protocol of such a test. It shows the integrated sensor-telemetry system response to glucose concentration increase to 200 mg d l - ' during the second day of continuous monitoring of 100 mg dl - ~ glucose. It can be seen from the Fig. 6 that during the alternate immersion of the device in 100 and 200 mg dl- ' glucose solutions the transmitted signal each time returns to the previous value. Comparison between the transmitted signal values of the cal,'~ration test of the integrated sensor-telemetry system (Fig. 5 L" ) with that obtained during the continuous glucose monitoring at 100 and at 200 mg d l - glucose shows the repeatability of the results within +2.5% of the main signal value.

4. Conclusions

An integrated glucose-monitoring system consisting of a glucose sensor, miniature potentiostat and FM signal transmitter has been developed. The integrated system was embedded in an epoxy cement or in a silastic matrix to ensure hermeticity. The device components at different stages of the construction were evaluated in vitro. A good correlation between the amperometric sensor signal and the demodulated telemetry signal has been demonstrated. To simulate implantation conditions, tests were conducted with the electronics encapsulated in plastic (epoxy resin or silicone elastomer matrix) and submerged in a 37 °C water bath. These tests verified that the integrated sensor-telemetry system is able to monitor glucose concentration adequately in vitro.

Long-term in vitro tests of the integrated system are currently in progress. The results collected will allow the commencement of in vivo experimentation in model dogs.

Acknowledgements

This research has been supported by a grant from the National Science Foundation and the Whitaker Foundation. The authors are grateful to Dr D. Williams, Mr J. Love and Mr S. Kuzmaul from Sandia National Laboratories, Albu- querque, for their technical help in the initial stage of this work. Encapsulation of the integrated units in epoxy cement was by courtesy of Guiton Co., Albuquerque. The authors are grateful to Professor A.G. Winfield (Training and Research Institute for Plastics, UNM) for his help in encapsulation of the units in silastic.

References

[I] K.F. Hanssen, H.-J. Bangstad, O. Brinchmann-Hansen, K. Dahl- Jorgensen, Blood glucose control and diabetic microvascular complications; long term effects of near normoglycaemia, Diabetic Medicine, 9 (1992) 697-705.

[2] S.R. Page and L Peacock, Blood glucose monitoring: does technology help?, Diabetic Medicine, 10 (1993) 793-801.

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[3] D. Gordon, C.G. Semple and K.R. Paterson, Do different frequencies of self monitoring of blood glucose influence control in type i diabetic patients?, Diabetic Medicine, 8 ( 1991 ) 679-682.

[4] R.R. Wing, D. Lamparski, S. Zaslow, J. BeLschart, J. Simiuerio and D. Becket, Frequency and accuracy of self monitoring of blood glucose in children: relationship to glycemic control. Diabetes Care. 8 ( 1985 ) 214-218.

[5] K. Waxman, D. Turner and T. Nguen. Implantable programmable insulin pumps for the treatment of diahetes, Arch. Surg., 127 (1992) 1032-1041.

[6] H. Buchwald and T.D. Thomas, Implantable pumps: present progress and anticipated future advances, Trans. Am. Soc. Artif. Internal Organs, 38 (1992) 772-777.

[7] M. Zoltobrocki, Insulin delivery by implantable pumps, Horm. Metab. Res. (Supl. Set.). 26 (1992) 140-145.

[ 8 ] D.C. leuHer, Overview of biomedical t:iemetry techniques, Eng. Med. Biol., 3 (1983) 17-24.

[9] M. Shichiri, N. Asasawa, Y. Yamasaki, R. Kawamori and H. Abe, Telemetry glucose monitoring device with needle glucose sensor: a useful tool for blood glucose monitoring in diabetic individuals, Diabetes Care, 9 (1986) 298-301.

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{ 12l S.L. Xie and E. Wilkins, Rechargeable glucose electrodes for long- term implantation, J. Biomed. Eng.. 13 ( 1991 ) 375-378.

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Bk ,g raph ies

Joseph Black is now an M.Sc. student in electrical engineering after leaving the US Air Force where he was a captain in an engineering command. His research interests are in the area o f telecommunications, radio-electronics and computer hardware.

Michael Wilkins (Ph.D. from University o f Illinois, 1969) is ,-urrently a professor o f electrical engineering at the Uni- v e r i t y o f New Mexico. His research interests include various electronics and computer science fields.

Plamen Atanasov (1962) is currently a research assistant professor at the University of New Mexico, Chemical and Nuclear Engineering Department. His research interests are in the field o f bio-electrocbemistry, enzyme electrodes and biosensors.

Ebtisam Wilkins received her Ph.D. in chemical engineering from the University o f Virginia (1976) . Since 1978 she has been a professor ofcbemica l engineering s t ~ e University o f New Mexico, heading the Biomedical Engineering P~:ogram in the Department of Chemical and Nuclear Engineering. Her research activities are in various fields o f solar-energy collection and storage, blotechnology of waste removal and development and applications ofbiosensors . Her biomedical investigations are concentrated on the development o f an implantable glucose sensor for an artificial pan- creas, intravascular sensors for glucose monitoring in whole blood and biocompatibili ty o f implantable materials.

integrated sensor-telemetry system for in vivo glucose monitoring

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