Biosensors & Bioelectronics5 (1990) 3746
Development of a Potentially Wearable Glucose Sensor for Patients with Diabetes Mellitus: Design and In-vitro Evaluation
A. J. M. Schoonen, F. J. Schmidt, H. Hasper, D. A. Verbrugge, R. G. Tiessen &C. F. Lerk Department
of Pharmaceutical Technology and Biopharmaceutics, Faculty of Pharmacy, University of Groningen, Ant. Deusinglaan 2,9713 AW Groningen, The Netherlands
(Received 2 February 1989; revised version received and accepted 15 May 1989)
ABSTRACT A potentially wearable glucose sensor was developed, consisting of an oxygen electrode as detector and a dynamic enzyme perfusion system as selector. The selector is a hollow fibre, which can be placed subcutaneously and dialyses glucose from tissue fluid. In this design the problems of enzyme instability and oxygen limitation might be circumvented. The sensor measures glucose reliably for over two weeks, provided a new 10 ml syringe containing a glucose oxidase solution is connected to the system each day. Key wor&: glucose sensor, glucose oxidase, enzyme perfusion, subcutaneous, oxygen electrode.
INTRODUCTION The use of insulin pumps in diabetic treatment is increasing, although they are not feedback controlled on blood glucose level. At present the insulin infusion rate of these pumps is controlled with the aid of ‘glucosticks’. This method requires frequent blood sampling from fingertips. A continuously measuring glucose sensor could improve the blood glucose management considerably with the prospect of decreasing long37
Biosensors & Bioelectronics09%5663/89/$03~50@ 1989Elsevier Science Publishers Ltd, England. Printed in Great Britain
38
A. J. M. Schoonen et al.
term complications of diabetes mellitus and increasing the comfort for diabetic patients. A reliable glucose sensor is vital for the successful introduction of a wearable artificial pancreas. There are several possibilities of measuring glucose (Turner & Pickup, 1985), but the most used method, because of its specificity, is an amperometric electrode based on the enzyme glucose oxida& (GOD). The use of the enzyme GOD is reported by many authors (kessler et al., 1984; Gough et al., 1986; Mtiller et al., 1986 and Shichiri et al., 1986); the instability of the enzyme, however, turns out to be one of the main problems for a reliable glucose sensor. Immobilizing the enzyme to prolong its stability is not sufficient. At body temperature, the activity of the enzyme still decreases rapidly. Besides, an immobilized enzyme layer on an electrode, together with other selective membranes considerably delays the sensor response. The glucose sensor developed was designed to control a wearable insulin infusion system by measuring the glucose level subcutaneously. It circumvents the instability of the enzyme by applying a dynamic enzyme perfusion system. The wearable glucose sensor, called a glucosensor, consists of a ‘selector’ part and a ‘detector’ part. The selector part is necessary to give the sensor its specificity for glucose in biological fluids. Two membranes and the enzyme GOD are used as selectors, while a very small oxygen electrode is used as the detector. The measuring principle of the glucosensor is the well-known reaction: GOD
Glucose + O2 j
Glucono-Glactone + H202
The electrode may be used both for measurement of O2 and H,02, depending on the polarity of the applied voltage. By changing one or more of the selectors (for example the enzyme), the system can be used for on-line measurement of numerous different molecules of interest.
METHODS The selector
The major part of the glucose selection procedure is carried out by a subcutaneous enzyme perfusion system. A hollow fibre (regenerated cellulose ester. Spectrum Medical Ind., Los Angeles, USA, o.d. 186 pm, i.d. 150pm, MWCO 9000 Dalton), connected on both sides to polyethylene tubes (i.d. O-4 mm, o.d. O-8 mm0 is placed subcutaneously
A wearable glucose sensor for diabetes mellitus patients
39
and perfused (Braun VI perfusor pump, Melsungen, FRG) with an enzyme solution (GOD; grade III lyophilized, 20-000 U, Boehringer Mannheim, Mannheim, FRG; concentration O-15 mg/ml in saline, 0.9% NaCl in H20). Saline and GOD were not made up in buffer. Glucose diffuses from the subcutaneous tissue into the hollow fibre where the reaction is catalysed by the enzyme. By connecting the polyethylene tube which comes out of the body to an oxygen electrode, the pOZ of the perfusion fluid can be measured. The amount of oxygen consumed in the reaction is related to the subcutaneous glucose concentration in a 1:l ratio, given by the reaction above. The dialysis step in the glucose sensor is a simple way to measure tissue glucose just outside the body. It is also a necessary step to obtain a dilution of the body glucose concentration. Saturated oxygen concentrations in water or body fluid are too low to complete the reaction with physiological glucose concentrations (oxygen limitation). The detector
The detector part of the glucosensor is a miniaturized Clark-type oxygen electrode (Clark & Lyons, 1962). A central platinum cathode is melted into glass and surrounded by a silver anode (two-electrode system). A 0.06 M K2HP04 solution is used as an electrolyte (Fig. 1). The electrode is completely covered with a teflon membrane, which is permeable only to gases. The specificity for oxygen is obtained using a standard laboratory model potentiostat, by applying a negative voltage of O-6 V to the working electrode (Pt) with respect to the reference electrode (Ag/AgCl) * The electrode reaction on the platinum tip is represented: 0,+2H,O+4e--4OHThe electrode reaction at the silver cathode is represented: 4Ag------+
4 Ag+ + 4e-
An advantage of the use of a teflon membrane is the almost complete separation between the perfusion fluid and the electrolytic solution. Interference of other constituents of the perfusion fluid on the electrode reaction is negligible. Some charactersistics of the oxygen electrode are: current (oxygen tension 20%) residual
current
lOCk120 nA 6-10 nA
A. J.
40
M. Schoonen
et al.
cell cavity
A 7.6 mm
flow chamber
Fig. 1. Oxygen electrode
and flow chamber.
The subcutaneous perfusion system is connected to the electrode by means of a flow chamber (Fig. 1). The perfusion fluid comes directly to the platinum tip of the oxygen electrode, where p02 is measured. The flow chamber was made as small as possible to avoid mixing and dilution effects, and to minimize the response time of the sensor. The flow chamber can be slipped over the electrode until the platinum tip is just above the opening and a symmetric flow regimen on the cathode is realized. Fluid from the flow chamber is caught in a little waste-bag. All in-vitro results were obtained at room temperature (2&25”C), which was kept as constant as possible. The hollow fibre was placed in unstirred glucose solutions over a range of concentrations. Stirring of the solutions (up to 400 r.p.m.) had no effect on sensor response. Initially, the response of the sensor was very low in freshly made glucose solutions. After equilibration of the glucose solutions for a 3-h period, the response remained unchanged. Apparently this is a mutarotation effect of glucose (Glucose Monohydrat, Merck, Darmstadt, FRG). Therefore all experiments carried out with glucose solutions were allowed to equilibrate for at least 4 h. RESULTS
AND DISCUSSION
Figure 2 shows a schematic sensor response curve with a sudden change in glucose concentration.
A wearable glucose sensor for diabetes mellitus patients
41
Fig. 2. Sensor response curve in vitro. At T = 0 change in glucose concentration.
TABLE 1
Sensor Performance In Vitro” Glucose concentration (m&W
Sensor output (n = 4) (n-4)
100 200 300
104.6 + 3.2’ 71-8 + 3.9 36-2 + 3.2
r: O-992 2 aHollow fibre length = 10 mm; perfusion fluid = GOD, O-15mg/ml in saline; perfusion rate = 0.6 ml/h; tube length between fibre and electrode = 1.5 cm; T 90% 3.1 f 0.4 min. bValues given are the means f SD.
Table 1 shows a sensor performance for a typical in-vitro configuration. The effect of enzyme perfusion rate on sensor output was observed. Figure 3 shows sensor output, as expected, was inversely proportional to enzyme perfusion rate. The same effect can be seen for fibre length, which was proportional to sensor output (see Fig. 4). It is obvious that the perfusion rate of the enzyme solution has to be kept as low as possible to minimize the daily use of enzyme. A low perfusion rate makes it possible to keep the syringe with enzyme solution small. Decreasing the perfusion rate is limited because of its negative effect on the response time. Although ultrafiltration of the hollow fibre solution might occur, leading to dilution of local tissue glucose, the effect, if present, is small and could not be measured in vitro. Microdialysis techniques are already used in in-vivo studies (Liinnroth et al., 1987). From these experiments it can be deduced that ultrafiltration does not affect the measurements. After optimization of these parameters a linear response was obtained from 0 to 400 mg/dl glucose (see Fig. 4).
A. J. M. Schoonen et al.
42
80-
ii ‘:
70-
i
60-
1 50-
40-
30-
20-
lo-
J Oo.2
03
1 0.4 l/perfusion
0.5 rate
06
(h/ml)
Fig. 3. Relationship
between enzyme perfusion rate and sensor output. A = 1000 mg/dl glucose, r = 0.996; 0 = 400 mg/dl glucose, r = O-993; -- -- = oxygen limitation.
Figure 5 shows a linear correlation between glucose concentration and the slope of the sensor response curves. This good correlation makes it possible to predict a steady state glucose concentration as soon as a sensor response occurs. Usually the time to reach 90% of the maximum value (T 90%) is about 3 min, and includes a lag time of less than 1 min if a sudden change in glucose concentration occurs. Thus it is only a matter of a few seconds after the lag time before an accurate prediction of the (static) glucose level can be made. An advantage of the glucosensor is the possibility of changing some of the important parameters. Response time, lag time, T 90%, current change per milligramme of glucose etc., can be manipulated by changing the enzyme perfusion rate, length and internal diameter (i.d.) of the hollow fibre and/or the length and i.d. of the connection tube between fibre and electrode. The oxygen electrode measures continuously for over 3 weeks without a basic drift in current (~2%). After this time the buffering capacity of the electrolyte becomes insufficient and a daily drift (more than 5%) occurs. Therefore the electrolyte should be refreshed after 2 weeks of operation. The perfusion system, as presented, is a first experimental approach to
A wearable glucose sensor for diabetes mellitus patients
43
I-
)-
t 0
Fig. 4. Dependency
concentrations.
100
glucose
200 concentratmn
300 (mgldll
400
of sensor output current on hollow fibre length at different glucose A = length lOmm, r = 0.997; 0 = length 5 mm, r = 0495.
if the method of this ‘dynamic’ glucose sensor works in vitro and in for 3 days at the most. Therefore, at present, functioning of the sensor without maintenance is limited to 3 days. Although it is easy to connect a new syringe, the aim of future work could be to design a glucose sensor with a closed loop perfusion system, in which the perfusion fluid is reused. With the use of an enzyme dosage system it should be possible to expand the lifetime of the sensor to at least a few weeks. It is well known that Clark-type oxygen electrodes are rather sensitive to changes in temperature. Diffusion of oxygen and solubility of oxygen in membranes and perfusion fluid are all parameters dependent on temperature. Generally an increase in temperature gives an increase in output current of Clark-type oxygen electrodes (Koryta, 1975). The mean temperature dependency of Clark-type oxygen electrodes is
check
vivo. A syringe full of GOD is stable at room temperature
A. J. M. Schoonen et al.
44
25
20
15 ? t ‘E .z H ii
10
5
0
100 ~lucore
200
concentration
300 Cmgldl)
400
Fii. 5. Relationship between glucose concentration and slope of the sensor response curves. A = hollow fibre length of 10 mm, r = O-995; 0 = hollow fibre length of 5 mm, r = O-996. Correlations between sensor output and the slope of the sensor response curves are 0496 and 0.982 respectively.
2% of the normal output current value with every degree Celsius (Sawyer et al., 1982). The oxygen electrode used in this work gave a linear increase in output current from 80 to 130 nA, with a temperature rise from 16 to 38°C (Fig. 6), resulting in a temperature dependency of 2el%/“C. Thus, temperature changes in the electrcde should be measured and compensated for electronically, if reliable in-vitro and in-vivo data are to be obtained. The perfusion fluid flowing through the hollow fibre can also exchange oxygen with the surrounding fluid. Minor changes in PO,! of the glucose solutions did not affect sensor functioning. Subcutaneous (fatty) tissue oxygen concentration varies between 95 and 98% of the maximum pOz in body fluid (large fluctuations in pO* only occur in muscles). So 02-
A wearable glucose sensor for diabetes mellitus patients
14
16
16
20
22
26
24
26
30
32
34
36
36 T ‘C
Fig. 6. Temperature
dependency
of the glucosensor;
l
= day 1; X = day 2.
exchange within the body is not likely to affect the final measured luminal pOZ. Further research, especially in viva, is needed to examine these parameters.
CONCLUSION The glucosensor described is an accurate and specific tool for measuring glucose concentration levels. The system circumvents the enzyme instability problem, as within a few seconds a syringe with fresh enzyme solution can be connected to the system. Having developed a reliable long-term functioning oxygen electrode, the lifetime of the glucosensor is expected to be at least two weeks. The only part of the sensor that has to be implanted subcutaneously in the body is the hollow fibre. An injection system to position the fibre correctly is now available, so currently the glucosensor is being tested in healthy volunteers and diabetic patients.
ACKNOWLEDGEMENT The authors wish to thank Mr E. H. Bruins, who designed an oxygen electrode on which our miniaturized oxygen electrode ,was based.
A. J.
46
M. Schoonen et al.
REFERENCES Clark, L. C. & Lyons, Ch. (1962). Electrode systems for continuous monitoring in cardiovascular surgery. Am. N. Y. Acud. Sci., 102,29-45. Gough, D. A., Armour, J. C., Lucisano, J. Y. & McKean, B. D. (1986). Short term in vivo operation of a glucose sensor. Trans. Am. Sot. Artif. Znfern. Organs, 32,14850. Kessler, M., Hoper, J., Volkholz, H.-J., Sailer, D. & Demling, L. (1984). A new glucose electrode for tissue measurements. Hepato-gasfroenterol., 31,285-7. Koryta, J. (1975). Enzyme Electrodes In Zon selective electrodes. Cambridge University Press, Cambridge, UK, pp. 174-7. Lonnroth, P., Jannson, P.-A. & Smith, U. (1987). A microdialysis method allowing characterization of intercellular water space in humans. Am. J. Physiol., 253, E228-31.
Mtiller, A., Abel, P. &Fischer, U. (1986). Continuously monitored subcutaneous glucose concentration by using implanted enzyme electrodes. Biomed. Biochem. Acta, 45(6), 769-77.
Sawyer, D. T., Chiericato, G., Jr, Angells, C. T., Nanni, E. J. & Tsuchlya, T. (1982). Effects of media and electrode materials on the electrochemical reduction of dioxygen. Anal. Chem., 54, 1720-4. Shichiri, M., Asakawa, N., Yamasaki, Y., Kawamori, R. & Abe, H. (1986). Telemetry glucose monitoring device with needle-type glucose sensor. A useful tool for blood glucose monitoring in diabetic individuals. Diabetes Cure, 9(3), 298-301.
Turner, A. P. F. & Pickup, J. C. (1985). Diabetes mellitus: biosensors research and management. Biosensors, 1,85-115.
for
Development of a Potentially Wearable Glucose Sensor for Patients with Diabetes Mellitus: Design and In-vitro Evaluation
A. J. M. Schoonen, F. J. Schmidt, H. Hasper, D. A. Verbrugge, R. G. Tiessen &C. F. Lerk Department
of Pharmaceutical Technology and Biopharmaceutics, Faculty of Pharmacy, University of Groningen, Ant. Deusinglaan 2,9713 AW Groningen, The Netherlands
(Received 2 February 1989; revised version received and accepted 15 May 1989)
ABSTRACT A potentially wearable glucose sensor was developed, consisting of an oxygen electrode as detector and a dynamic enzyme perfusion system as selector. The selector is a hollow fibre, which can be placed subcutaneously and dialyses glucose from tissue fluid. In this design the problems of enzyme instability and oxygen limitation might be circumvented. The sensor measures glucose reliably for over two weeks, provided a new 10 ml syringe containing a glucose oxidase solution is connected to the system each day. Key wor&: glucose sensor, glucose oxidase, enzyme perfusion, subcutaneous, oxygen electrode.
INTRODUCTION The use of insulin pumps in diabetic treatment is increasing, although they are not feedback controlled on blood glucose level. At present the insulin infusion rate of these pumps is controlled with the aid of ‘glucosticks’. This method requires frequent blood sampling from fingertips. A continuously measuring glucose sensor could improve the blood glucose management considerably with the prospect of decreasing long37
Biosensors & Bioelectronics09%5663/89/$03~50@ 1989Elsevier Science Publishers Ltd, England. Printed in Great Britain
38
A. J. M. Schoonen et al.
term complications of diabetes mellitus and increasing the comfort for diabetic patients. A reliable glucose sensor is vital for the successful introduction of a wearable artificial pancreas. There are several possibilities of measuring glucose (Turner & Pickup, 1985), but the most used method, because of its specificity, is an amperometric electrode based on the enzyme glucose oxida& (GOD). The use of the enzyme GOD is reported by many authors (kessler et al., 1984; Gough et al., 1986; Mtiller et al., 1986 and Shichiri et al., 1986); the instability of the enzyme, however, turns out to be one of the main problems for a reliable glucose sensor. Immobilizing the enzyme to prolong its stability is not sufficient. At body temperature, the activity of the enzyme still decreases rapidly. Besides, an immobilized enzyme layer on an electrode, together with other selective membranes considerably delays the sensor response. The glucose sensor developed was designed to control a wearable insulin infusion system by measuring the glucose level subcutaneously. It circumvents the instability of the enzyme by applying a dynamic enzyme perfusion system. The wearable glucose sensor, called a glucosensor, consists of a ‘selector’ part and a ‘detector’ part. The selector part is necessary to give the sensor its specificity for glucose in biological fluids. Two membranes and the enzyme GOD are used as selectors, while a very small oxygen electrode is used as the detector. The measuring principle of the glucosensor is the well-known reaction: GOD
Glucose + O2 j
Glucono-Glactone + H202
The electrode may be used both for measurement of O2 and H,02, depending on the polarity of the applied voltage. By changing one or more of the selectors (for example the enzyme), the system can be used for on-line measurement of numerous different molecules of interest.
METHODS The selector
The major part of the glucose selection procedure is carried out by a subcutaneous enzyme perfusion system. A hollow fibre (regenerated cellulose ester. Spectrum Medical Ind., Los Angeles, USA, o.d. 186 pm, i.d. 150pm, MWCO 9000 Dalton), connected on both sides to polyethylene tubes (i.d. O-4 mm, o.d. O-8 mm0 is placed subcutaneously
A wearable glucose sensor for diabetes mellitus patients
39
and perfused (Braun VI perfusor pump, Melsungen, FRG) with an enzyme solution (GOD; grade III lyophilized, 20-000 U, Boehringer Mannheim, Mannheim, FRG; concentration O-15 mg/ml in saline, 0.9% NaCl in H20). Saline and GOD were not made up in buffer. Glucose diffuses from the subcutaneous tissue into the hollow fibre where the reaction is catalysed by the enzyme. By connecting the polyethylene tube which comes out of the body to an oxygen electrode, the pOZ of the perfusion fluid can be measured. The amount of oxygen consumed in the reaction is related to the subcutaneous glucose concentration in a 1:l ratio, given by the reaction above. The dialysis step in the glucose sensor is a simple way to measure tissue glucose just outside the body. It is also a necessary step to obtain a dilution of the body glucose concentration. Saturated oxygen concentrations in water or body fluid are too low to complete the reaction with physiological glucose concentrations (oxygen limitation). The detector
The detector part of the glucosensor is a miniaturized Clark-type oxygen electrode (Clark & Lyons, 1962). A central platinum cathode is melted into glass and surrounded by a silver anode (two-electrode system). A 0.06 M K2HP04 solution is used as an electrolyte (Fig. 1). The electrode is completely covered with a teflon membrane, which is permeable only to gases. The specificity for oxygen is obtained using a standard laboratory model potentiostat, by applying a negative voltage of O-6 V to the working electrode (Pt) with respect to the reference electrode (Ag/AgCl) * The electrode reaction on the platinum tip is represented: 0,+2H,O+4e--4OHThe electrode reaction at the silver cathode is represented: 4Ag------+
4 Ag+ + 4e-
An advantage of the use of a teflon membrane is the almost complete separation between the perfusion fluid and the electrolytic solution. Interference of other constituents of the perfusion fluid on the electrode reaction is negligible. Some charactersistics of the oxygen electrode are: current (oxygen tension 20%) residual
current
lOCk120 nA 6-10 nA
A. J.
40
M. Schoonen
et al.
cell cavity
A 7.6 mm
flow chamber
Fig. 1. Oxygen electrode
and flow chamber.
The subcutaneous perfusion system is connected to the electrode by means of a flow chamber (Fig. 1). The perfusion fluid comes directly to the platinum tip of the oxygen electrode, where p02 is measured. The flow chamber was made as small as possible to avoid mixing and dilution effects, and to minimize the response time of the sensor. The flow chamber can be slipped over the electrode until the platinum tip is just above the opening and a symmetric flow regimen on the cathode is realized. Fluid from the flow chamber is caught in a little waste-bag. All in-vitro results were obtained at room temperature (2&25”C), which was kept as constant as possible. The hollow fibre was placed in unstirred glucose solutions over a range of concentrations. Stirring of the solutions (up to 400 r.p.m.) had no effect on sensor response. Initially, the response of the sensor was very low in freshly made glucose solutions. After equilibration of the glucose solutions for a 3-h period, the response remained unchanged. Apparently this is a mutarotation effect of glucose (Glucose Monohydrat, Merck, Darmstadt, FRG). Therefore all experiments carried out with glucose solutions were allowed to equilibrate for at least 4 h. RESULTS
AND DISCUSSION
Figure 2 shows a schematic sensor response curve with a sudden change in glucose concentration.
A wearable glucose sensor for diabetes mellitus patients
41
Fig. 2. Sensor response curve in vitro. At T = 0 change in glucose concentration.
TABLE 1
Sensor Performance In Vitro” Glucose concentration (m&W
Sensor output (n = 4) (n-4)
100 200 300
104.6 + 3.2’ 71-8 + 3.9 36-2 + 3.2
r: O-992 2 aHollow fibre length = 10 mm; perfusion fluid = GOD, O-15mg/ml in saline; perfusion rate = 0.6 ml/h; tube length between fibre and electrode = 1.5 cm; T 90% 3.1 f 0.4 min. bValues given are the means f SD.
Table 1 shows a sensor performance for a typical in-vitro configuration. The effect of enzyme perfusion rate on sensor output was observed. Figure 3 shows sensor output, as expected, was inversely proportional to enzyme perfusion rate. The same effect can be seen for fibre length, which was proportional to sensor output (see Fig. 4). It is obvious that the perfusion rate of the enzyme solution has to be kept as low as possible to minimize the daily use of enzyme. A low perfusion rate makes it possible to keep the syringe with enzyme solution small. Decreasing the perfusion rate is limited because of its negative effect on the response time. Although ultrafiltration of the hollow fibre solution might occur, leading to dilution of local tissue glucose, the effect, if present, is small and could not be measured in vitro. Microdialysis techniques are already used in in-vivo studies (Liinnroth et al., 1987). From these experiments it can be deduced that ultrafiltration does not affect the measurements. After optimization of these parameters a linear response was obtained from 0 to 400 mg/dl glucose (see Fig. 4).
A. J. M. Schoonen et al.
42
80-
ii ‘:
70-
i
60-
1 50-
40-
30-
20-
lo-
J Oo.2
03
1 0.4 l/perfusion
0.5 rate
06
(h/ml)
Fig. 3. Relationship
between enzyme perfusion rate and sensor output. A = 1000 mg/dl glucose, r = 0.996; 0 = 400 mg/dl glucose, r = O-993; -- -- = oxygen limitation.
Figure 5 shows a linear correlation between glucose concentration and the slope of the sensor response curves. This good correlation makes it possible to predict a steady state glucose concentration as soon as a sensor response occurs. Usually the time to reach 90% of the maximum value (T 90%) is about 3 min, and includes a lag time of less than 1 min if a sudden change in glucose concentration occurs. Thus it is only a matter of a few seconds after the lag time before an accurate prediction of the (static) glucose level can be made. An advantage of the glucosensor is the possibility of changing some of the important parameters. Response time, lag time, T 90%, current change per milligramme of glucose etc., can be manipulated by changing the enzyme perfusion rate, length and internal diameter (i.d.) of the hollow fibre and/or the length and i.d. of the connection tube between fibre and electrode. The oxygen electrode measures continuously for over 3 weeks without a basic drift in current (~2%). After this time the buffering capacity of the electrolyte becomes insufficient and a daily drift (more than 5%) occurs. Therefore the electrolyte should be refreshed after 2 weeks of operation. The perfusion system, as presented, is a first experimental approach to
A wearable glucose sensor for diabetes mellitus patients
43
I-
)-
t 0
Fig. 4. Dependency
concentrations.
100
glucose
200 concentratmn
300 (mgldll
400
of sensor output current on hollow fibre length at different glucose A = length lOmm, r = 0.997; 0 = length 5 mm, r = 0495.
if the method of this ‘dynamic’ glucose sensor works in vitro and in for 3 days at the most. Therefore, at present, functioning of the sensor without maintenance is limited to 3 days. Although it is easy to connect a new syringe, the aim of future work could be to design a glucose sensor with a closed loop perfusion system, in which the perfusion fluid is reused. With the use of an enzyme dosage system it should be possible to expand the lifetime of the sensor to at least a few weeks. It is well known that Clark-type oxygen electrodes are rather sensitive to changes in temperature. Diffusion of oxygen and solubility of oxygen in membranes and perfusion fluid are all parameters dependent on temperature. Generally an increase in temperature gives an increase in output current of Clark-type oxygen electrodes (Koryta, 1975). The mean temperature dependency of Clark-type oxygen electrodes is
check
vivo. A syringe full of GOD is stable at room temperature
A. J. M. Schoonen et al.
44
25
20
15 ? t ‘E .z H ii
10
5
0
100 ~lucore
200
concentration
300 Cmgldl)
400
Fii. 5. Relationship between glucose concentration and slope of the sensor response curves. A = hollow fibre length of 10 mm, r = O-995; 0 = hollow fibre length of 5 mm, r = O-996. Correlations between sensor output and the slope of the sensor response curves are 0496 and 0.982 respectively.
2% of the normal output current value with every degree Celsius (Sawyer et al., 1982). The oxygen electrode used in this work gave a linear increase in output current from 80 to 130 nA, with a temperature rise from 16 to 38°C (Fig. 6), resulting in a temperature dependency of 2el%/“C. Thus, temperature changes in the electrcde should be measured and compensated for electronically, if reliable in-vitro and in-vivo data are to be obtained. The perfusion fluid flowing through the hollow fibre can also exchange oxygen with the surrounding fluid. Minor changes in PO,! of the glucose solutions did not affect sensor functioning. Subcutaneous (fatty) tissue oxygen concentration varies between 95 and 98% of the maximum pOz in body fluid (large fluctuations in pO* only occur in muscles). So 02-
A wearable glucose sensor for diabetes mellitus patients
14
16
16
20
22
26
24
26
30
32
34
36
36 T ‘C
Fig. 6. Temperature
dependency
of the glucosensor;
l
= day 1; X = day 2.
exchange within the body is not likely to affect the final measured luminal pOZ. Further research, especially in viva, is needed to examine these parameters.
CONCLUSION The glucosensor described is an accurate and specific tool for measuring glucose concentration levels. The system circumvents the enzyme instability problem, as within a few seconds a syringe with fresh enzyme solution can be connected to the system. Having developed a reliable long-term functioning oxygen electrode, the lifetime of the glucosensor is expected to be at least two weeks. The only part of the sensor that has to be implanted subcutaneously in the body is the hollow fibre. An injection system to position the fibre correctly is now available, so currently the glucosensor is being tested in healthy volunteers and diabetic patients.
ACKNOWLEDGEMENT The authors wish to thank Mr E. H. Bruins, who designed an oxygen electrode on which our miniaturized oxygen electrode ,was based.
A. J.
46
M. Schoonen et al.
REFERENCES Clark, L. C. & Lyons, Ch. (1962). Electrode systems for continuous monitoring in cardiovascular surgery. Am. N. Y. Acud. Sci., 102,29-45. Gough, D. A., Armour, J. C., Lucisano, J. Y. & McKean, B. D. (1986). Short term in vivo operation of a glucose sensor. Trans. Am. Sot. Artif. Znfern. Organs, 32,14850. Kessler, M., Hoper, J., Volkholz, H.-J., Sailer, D. & Demling, L. (1984). A new glucose electrode for tissue measurements. Hepato-gasfroenterol., 31,285-7. Koryta, J. (1975). Enzyme Electrodes In Zon selective electrodes. Cambridge University Press, Cambridge, UK, pp. 174-7. Lonnroth, P., Jannson, P.-A. & Smith, U. (1987). A microdialysis method allowing characterization of intercellular water space in humans. Am. J. Physiol., 253, E228-31.
Mtiller, A., Abel, P. &Fischer, U. (1986). Continuously monitored subcutaneous glucose concentration by using implanted enzyme electrodes. Biomed. Biochem. Acta, 45(6), 769-77.
Sawyer, D. T., Chiericato, G., Jr, Angells, C. T., Nanni, E. J. & Tsuchlya, T. (1982). Effects of media and electrode materials on the electrochemical reduction of dioxygen. Anal. Chem., 54, 1720-4. Shichiri, M., Asakawa, N., Yamasaki, Y., Kawamori, R. & Abe, H. (1986). Telemetry glucose monitoring device with needle-type glucose sensor. A useful tool for blood glucose monitoring in diabetic individuals. Diabetes Cure, 9(3), 298-301.
Turner, A. P. F. & Pickup, J. C. (1985). Diabetes mellitus: biosensors research and management. Biosensors, 1,85-115.
for
