Biosensors & Bioelectronics5 (1990) 1-12
Thermoelectric Enzyme Sensor for Measuring Blood Glucose
Michael J. Muehlbauer,* Eric J. Guilbeau, Bruce C. Towe & Tedd A. Brandoi Arizona State University, Departmentof Chemical, Bio and MaterialsEngineering, Tempe, Arizona 85227,USA (Received 17February1989;accepted14 April 1989)
ABSTRACT A new calorimetric sensor has been developed which employs a thin-film thermopile in association with an immobilized enzyme. The thermopile detects the minute temperature rise that occurs when a specific chemical substrate is catalyzed by the enzyme. A prototype sensor is described which generates an equivalent proportional voltage response to glucose concentrations present in either buffer solution or blood. These sensors have remained useful for up to 18 days when operated intermittently for measuring glucose in buffer solutions, or for up to 4 days when operated continuously. When implanted inside cardiovascular shunts on anesthetized dogs, the sensors responded appropriately to changes in the blood glucose concentration. Key words: blood glucose, calorimetric, catalase, catheter, enzyme, glucose oxidase, implantation, thermoelectric, thermopile, stability.
INTRODUCTION The concept of a general calorimetric sensor for detecting biochemical reagents by virtue of their enzymatic enthalpy of reaction is attractive in terms of its versatility and universality. All previous sensors of this type *Presentaddress:Cytogam, Inc., 3498N San MarcosPlace, Suite 7, Chandler,Arizona 85224, USA. Biosensors & Bioelectronics 09%5663/89/$03-50
Ltd, England. Printed in Great Britain
0 1989 Elsevier Science Publishers
M. J. Muehlbauer,
2
E. J. Guilbeau, B. C. Towe, T. A. Brandon
have been fabricated with thermistors (Weaver et al., 1976; Tran-Minh &Vallin, 1978; Rich et al., 1979; Fulton & Cooney, 1980). While providing adequate sensitivity for the measurement of the extremely small temperature differences produced, thermistors have suffered in regards to their poor signal-to-noise ratio and problems associated with their necessity for an external voltage bias. An alternative calorimetric sensor has recently been advanced based on the thermoelectric principle (Guilbeau et al., 1988). With this device, enzyme is immobilized directly over one set of thermoelectric junctions on a thin-film thermopile. In response to a specific chemical substrate, the enzyme generates heat thereby establishing a temperature gradient across the thermopile. A voltage output results by virtue of the thermoelectric Seebeck effect whose magnitude is proportibnal to the concentration of substrate in the bulk solution. Initial experiments with thermoelectric enzyme sensors employed the enzyme glucose oxidase (GOD) coupled to the sensor surface to catalyze the reaction GOD
Glucose + O2
-+
Gluconic acid + H202 + 79 kJ/mol
Clearly, oxygen acts as a co-reactant and must remain in excess in order for the sensor response to be proportional to the concentration of glucose. Towards this purpose, a second enzyme, catalase, is co-immobilized to catalyze the breakdown of hydrogen peroxide created in the first reaction Catalase H202
-
H20 + l/2 O2 + 100 kJ/mol
The reaction regenerates oxygen and acts to extend the range of response of the sensor to glucose, and further provides an additional enthalpy of reaction which more than doubles the sensitivity. Previous results with the thermoelectric enzyme sensor indicated the dependence of its voltage output upon the amount of enzyme immobilized, and also upon the flow rate and temperature (Muehlbauer et al., 1989). Its response time, however, was extremely rapid (less than 6 s). Moreover, in comparison to enzyme-coupled thermistor devices, the signal-to-noise ratio and baseline stability were remarkably enhanced. Based on these characteristics, an investigation into the possibility of employing the thermoelectric enzyme sensor for measuring glucose in blood for applications in biomedicine or in clinical analysis was undertaken. The following reports the results of experiments for determining the response of the sensor in whole blood both in vitro and in a physiological environment within a cardiovascular shunt on an anesthetized dog. To
Thermoelectric enzyme sensor for measuring blood glucose
3
ascertain the applicability of the sensor for long-term glucose monitoring, we tested its stability in phosphate buffer solution both on a continuous basis and intermittently over a period of several days.
SENSOR FABRICATION The thermoelectric enzyme sensor was constructed in a probe form as illustrated in Fig. 1. Its fabrication has been described previously (Muehlbauer et al., 1989). Briefly, the thermopile consists of 50 couples of bismuth and antimony metal evaporated onto a Mylar (polyethylene terephthalate) insulating support. Individual thermopiles were cut and
Enzyme-coated thermopile
Shielded wire
Polyethylene
tube
EPOXY
\ Ground wire
Fig. 1. Thermoelectric
enzyme sensor.
rolled into cylindrical form for mounting at the tip of a 3-mm polyethylene catheter. The metals of the thermopile were positioned facing the interior of the cylinder and were bonded at their contact pads with shielded electrical wire. This wire extended through the catheter for external connection with a microvoltmeter. Polyurethane foam was sprayed through the interior of the catheter to retain the Mylar in a cylindrical shape. The seams of the Mylar were sealed with epoxy. A modification of the sensor for adaptation to blood measurements was the inclusion of additional electrical shielding over the thermopile. For this purpose, a thread of the wire-gauze shield covering the contact wire was positioned on the exterior surface of the catheter. The external surface of the Mylar insulation opposite to the thermopile was coated with a thin layer of silver conductive paint and connected with the wire shield. The shield was then appropriately grounded to reduce electrical interference effects. Enzyme was immobilized as previously described by first dipcoating the sensor in a solution of segmented polyurethane and then cross-linking a solution of GOD, catalase, and albumin with glutaraldehyde.
4
M. J. Muehlbauer, E. J. Guilbeau, B. C. Towe, T. A. Brandon
EXPERIMENTAL
METHODS
In-vitro experiments For the in-vitro experiments, sensors were tested in either phosphate buffer solution (pH 7.4) or in heparinized canine blood. The canine blood was exsanguinated from anesthetized dogs at the termination of in-vivo experiments and was reduced in glucose concentration by prior injection of insulin. A flow-loop apparatus was employed in which a peristaltic blood pump drew from a reservoir containing approximately 1 litre of either buffer solution or blood. The fluid passed through a loop of Tygon tubing inside which the sensor was inserted through one convergent leg of a Y-connector. Fluid flowed from the reservoir past the sensor in the same direction as the sensor tip was pointed. Prior to its return to the reservoir, the fluid was saturated with pure oxygen by passage through a bubble oxygenator. Voltage measurements from the sensor were obtained with a HewlettPackard null voltmeter and were measured relative to the zero baseline voltage of the amplifier. The voltage response to increasing concentrations of glucose was recorded following injections of a concentrated solution of glucose into the fluid reservoir. The concentration of glucose within the reservoir was measured after approximately 5 min of mixing by analyzing a sample on a Kodak Ektachem analyzer (Eastman Kodak Co., Rochester, New York, USA). Shunt implantation experiments For the purpose of testing the sensor in a physiological environment, a cardiovascular shunt was devised. The shunt consisted of a length of Tygon tubing for placement between the femoral artery and vein on opposite legs of anesthetized dogs. The sensor was again positioned in the flow by insertion through an in-line Y-connector. Separate Luer-Lok ports were also included in the shunt for sample withdrawal and for the insertion of transducers for temperature and pressure monitoring. Greyhound dogs (23-38 kg) were anesthetized intramuscularly with Ketamine Plus (11 mg ketamine, 2.2 mg xylazine, O-04 mg atropine per kg dog weight; Bristol Laboratories, Syracuse, New York, USA) followed by intravenous sodium pentobarbital. The shunt was placed following the surgical cutdown of the opposing femoral artery and vein, and the sensor was positioned in the direction of blood flow. The dog was respirated with pure oxygen with a Harvard positive volume respirator (Harvard Apparatus, South Natick, Massachusetts, USA) to maintain a blood oxygen
Thermoelectric enzyme sensor for measuring blood glucose
5
tension of roughly 400 mm Hg. Blood pH was maintained with injections of sodium bicarbonate. To ascertain the stability of the sensor in this physiological environment, the voltage output was recorded continuously and compared against glucose concentration measurements obtained from the Kodak analyzer on withdrawn blood samples. These samples were withdrawn every 10 or 15 min and were additionally analyzed for pH and blood gases with a Radiometer ABL2 analyzer (Radiometer Copenhagen, Copenhagen, Denmark). The response of the sensor to glucose was noted following the injection of a dose of insulin (Eli Lilly, Indianapolis, Indiana, USA; 1 U per kg dog weight). As a further test of stability, sensors were tested for their response to glucose in phosphate buffer solution both prior to and following the shunt experiments.
RESULTS Long-term stability tests
These tests were performed to ascertain the ability of the thermoelectric sensors to respond to glucose over a period of many days. Sensors were stored in buffer solutions at pH 7.4 and were periodically removed to test their response to glucose. Figure 2 displays the results for one particular sensor stored and tested in phosphate buffer at room temperature for a period of 18 days. Tests were performed with an oxygen saturated solution in which the concentration of glucose was progressively increased. The sensitivity to glucose for ADay1 *Days 2-4 *Days 6-l 1 n Day
4.54
13
VDays
16-18
4.
0
Fii.
50
100 150 GLUCOSE CONCENTRATION,
200 mg %
250
r 300
2. Long-term stability of thermoelectric enzyme glucose sensor stored and tested in phosphate buffer solution at room temperature.
6
M. J. Muehlbauer,
E. J. Guilbeau,
B. C. Towe, T. A. Brandon
this sensor decreased from an initial value of 15 nV/mg% on the first day of testing to a value of 6 nV/mg% on the eighteenth day. The range of linear response also declined from an initial value of 329 mg% to a final concentration of 140 mg%. The sensor failed to exhibit any response to glucose on the twentieth day of testing. In total, 6 individual sensors were tested for their long-term response to glucose. Three others were tested at room temperature-2 after storage at room temperature, a third with refrigerator storage. The remaining 2 sensors were stored at 37°C and tested at both room temperature and 37°C. These last sensors showed a larger response to glucose at the higher temperature, as expected based on the increased rate of glucose reaction. Typical characteristics for all sensors were a continual decline in both sensitivity and range of linear response. In general, the sensors responded to glucose for a longer period when stored at the lower temperatures. Sensors stored at room temperature remained useful for measuring glucose for 7,9, and 18 days; the sensor with refrigerator storage for 24 days; and those at 37°C for 5 and 6 days. The sensors typically failed by electrical failure, as characterized by a dramatic increase in signal noise which was believed to result from water permeation into the sensor core. The sensor represented in Fig. 2, however, did not exhibit this behavior and was believed to fail as a result of immobilized enzyme deactivation, as verified by an enzyme assay. Continuous
stability tests
Another set of stability tests were performed in which sensors were tested for their response to glucose during continuous operation. Here, each sensor was kept in a continuously flowing stream of buffer solution at room temperature and exposed to a constant concentration of glucose, typically 60 mg%. In all, 4 sensors were tested in this manner. In each case, the voltage output declined steadily over time from its initial value at first exposure to glucose. The normalized rate of decline, defined as the rate of voltage decline in V h-’ divided by the initial voltage response, was 0.0098 If:0.0030 h-‘. Three of the sensors appeared to fail by electrical failure and remained useful for periods of 30, 31, and 78 h. The fourth sensor remained useful for 97 h at which time its voltage response to glucose was reduced to zero, an apparent result of enzyme deactivation. In-vitro response to glucose in blood
Five different sensors were tested for their response to glucose in whole blood. The best results are displayed in Fig. 3. Here, the response to
Thermoelectric
enzyme sensor for measuring blood glucose
0 Phosphate buffer 0 Dog blood
solution
12+
c
,,,.,.,.,~,.,~r 0
Fig. 3. Comparison
50
100
200 250 150 OLUCOSE CONCENTRATION,
300 rng 96
350
400
450
of the response of the thermoelectric enzyme glucose sensor in blood and in phosphate buffer solution.
glucose in blood is compared to the response in phosphate buffer solution both before and after the blood measurements. The sensitivity to glucose in the buffer solution exhibited no change following the tests in blood, and remained stable at a value of 36 f 1 nV/mg%. The sensitivity in blood was 34 + 2 nV/mg%. The combined sensitivity to glucose was 35 + 1 nV/mg% and the range of linear response was 328 mg%. Other sensors tested for their response in whole blood often revealed a slightly different zero offset than they did when operated in buffer solution. In all cases, however, the sensitivity to glucose remained equivalent in the two solutions. The average sensitivity to glucose in blood was 42 k 10 nV/mg% compared to an average of 44 + 10 nV/mg% in phosphate buffer solution. Response to glucose in cardiovascular
shunts
Three different sensors were tested for their response to glucose when placed in cardiovascular shunts on anesthetized dogs. Figure 4 represents the response of one of these sensors which remained in the shunt for a period of 4 h. Here, the scale for the sensor output is converted into a glucose concentration scale with the aid of a linear regression calibration between the analyzer glucose measurements and the sensor voltage. Analyzer measurements of glucose performed on withdrawn blood samples are indicated in the figure as individual points. The calibrated glucose sensitivity was 47 nV/mg%. In Fig. 4, the response of the sensor following an injection of insulin into the dog is exhibited. The analyzer measure-
8
M. J. Muehlbauer, E. J. Guilbeau, B. C. Towe, T. A. Brandon INSULIN
IOmln
Fig. 4. Continuous
cardiovascular
output from a thermoelectric enzyme glucose sensor implanted in a shunt, indicating the response following injection of insulin. Vertical voltage scale is converted into concentration units.
Sensor
=
Glucose
I .Ol
Ektachem
Glucose
-
1.6
r2=
0.663
240. 220.
.
200. 180.
0
20
Fig. 5. Correlation
40
60
80 EKTACHEM
120 100 GLUCOSE, mg%
140
160
180
200
of glucose concentrations measured by the thermoelectric glucose sensor and by the Kodak Ektachem analyzer.
enzyme
ments indicate a decrease in glucose concentration from 120 to 30 mg% correlated with a decline in sensor voltage of approximately 4 pV. For all 3 shunt-implanted sensors combined, the average sensitivity to glucose measured in the blood stream was 40 + 7 nV/mg%. The sensitivity to glucose in phosphate buffer solution was comparable and dropped an average of 7% following the shunt implantation experiments. Figure 5 indicates the degree of correlation between the measurements of glucose concentration made with the 3 shunt-implanted sensors and those made on withdrawn blood samples with the Kodak analyzer. Because of the manner in which the sensors were calibrated, linear regression between the two measured concentrations of glucose produced a line with nearly unity slope. The value of ?, however, is significant in indicating the
Thermoelectric enzyme sensor for measuring blood glucose
9
uncertainty in the sensor measurement. As a result of this uncertainty, one particular value of glucose concentration as measured with the calibrated sensor was less than zero.
DISCUSSION The decline in sensor response to glucose over a period of many days as illustrated in Fig. 2 was typical for all sensors tested. In general, the initial linear sensitivity steadily declined as did the range of linear response, defined as the concentration of glucose at which the sensor output diverges from linearity. The former result was not surprising based on the rationale that enzyme deactivation reduced the rate of reaction and thereby the heat which the sensor detects. The second result was unexpected, however, based on simulations which suggested that a reduction in GOD activity should extend the range of linear response by allowing the reaction of glucose to be less diffusion limited (Muehlbauer et al., 1989). Apparently, the co-immobilized enzyme, catalase, was not present in a sufficient quantity to completely consume the hydrogen peroxide generated in the GOD-catalyzed reaction. Hydrogen peroxide causes deactivation of both enzymes (Buchholz & Godelmann, 1978; Tse & Gough, 1987) and is preferably eliminated by providing sufficient excess of catalase. However, as separate experiments with co-immobilized membranes of GOD and catalase illustrate, even an 80-fold excess in catalase actvity was not sufficient to achieve maximum enzyme stability (Ireland, 1988). The indication is that some hydrogen peroxide remains unconsumed. Without complete consumption of hydrogen peroxide, a stoichiometric ratio of oxygen to glucose consumption of more than 1:2 results. Furthermore, since catalase deactivates faster than GOD, the ratio is expected to increase over time. Consequently, the range of linear response for the sensor is reduced as oxygen becomes the limiting reactant at increasingly lower concentrations of glucose. Likewise, the sensitivity is further reduced as less heat is generated in the coupled catalase reaction. These observations are consistent with the results of the long-term stability tests. As with the long-term stability tests, sensors tested on a continuous basis showed a steady although more rapid decline in sensitivity. For both tests and in most cases, the sensors appeared to fail as a result of water permeation into the thin-film element of the device. This occurred despite the excellent hydrophobicity of the Mylar and the presence of the epoxy sealant. Upon failure, sensors often revealed an abrupt decrease in
10
M. J. Muehlbauer,
E. J. Guilbeau,
B. C. Towe, T. A. Brandon
signal-to-noise ratio and an inability to respond to glucose even though a subsequent enzyme assay revealed the presence of enzyme activity. These sensors often exhibited an increase in resistance, although not always to an infinite value. Other sensors appeared to fail by virtue of enzyme deactivation as indicated by a loss of assayed enzyme activity and a maintenance of good electrical characteristics. Upon initiation of the experiments for testing the sensor performance in blood both in vitro and in the shunt, a noise component in the signal was immediately discovered which was not present for the tests in buffer solution. The additional noise was partially attributed to thermal interference, but was also an apparent interaction of the sensor with the blood. It was postulated that electrical interference from streaming currents caused by the adherence of charged proteins was the cause. The addition of an electrical shield opposite the thin-film thermopile improved the signal output from the sensor for all experiments in which the sensor was immersed in blood. As indicated in the in-vitro blood experiments, the sensor response to glucose generally exhibited the same sensitivity whether the sensor was operated in blood or in buffer solution. This equivalence in sensitivity, as Fig. 3 typifies, was found when the concentration of glucose in blood was measured as the plasma value. The analytical measurements for glucose in blood in this figure were measured by spinning the samples in a centrifuge prior to determination on the Kodak Ektachem analyzer. The indication, therefore, is that the sensor appeared to respond to the plasma concentration of glucose despite its operation in whole blood. This is consistent with the consideration that the sensor operates in a steady-state mode and draws its glucose for measurement from the plasma. Such a relatively small amount is consumed that glucose from the blood cells is not extracted. The sensor response in the cardiovascular shunt was acted upon by considerable interference. Among the probable sources were thermal gradients, piezoelectric currents caused by the blood pressure, and bioelectric currents. The comparison between the analytical and the sensor measurements of glucose, as Fig. 4 illustrates, was obscured by noise, sensor voltage offset, and a minimal sensitivity to glucose for the thermoelectric sensor. Despite these complications, however, a discernible response to glucose was detected. The response following a decline in glucose concentration as initiated by insulin administration was consistent with expectation. In addition, although the sensor performed somewhat erratically in the shunt environment, the exposure did not appear to produce any serious deleterious effects. The sensor performed equally well both before and after implantation when operated in buffer solution. The reduction in sensitiv-
Thermoelectric enzyme sensor for measuring blood glucose
11
ity following implantation was no greater than that found for sensors operated continuously in buffer solutions. The thermoelectric enzyme glucose sensor represents the first calorimetric device to be implanted and operated in a continuous mode in the blood stream. Realistically, the sensor cannot compete with polarographic or optical enzyme sensors for similar applications. Disadvantages include its present size and reduced sensitivity. These experiments indicate, however, that anticipated problems associated with thermal and electrical interference and stability can be overcome. In truth, the signal-to-noise ratio of the thermoelectric sensor is comparable to that of other implanted glucose sensors (Shichiri et al., 1984, 1986). Its improved response time and adaptability to other enzyme-substrate systems therefore make it attractive for analytical purposes, if not for biomedical implantation. ACKNOWLEDGEMENTS This research was supported by grants from the Whitaker Foundation the Arizona Disease Control Research Commission.
and
REFERENCES Buchholz, K. & Godelmann, B. (1978). Macrokinetics and operational stability of immobilized glucose oxidase and catalase. Biotech. Bioeng., XX, 1201-20. Fulton, S. P. & Cooney, C. L. (1980). Thermal enzyme probe with differential temperature measurements in a laminar flow-through cell. Anal. Chem., 52, 505-8.
Guilbeau, E. J., Towe, B. C. & Muehlbauer, M. J. (1988). A potentially implantable thermoelectric sensor for measurement of glucose. Trans. Am. Sot. Artif. Intern. Organs, XxX111, 329-35. Ireland, T. W. (1988). In vitro and in vivo stability of co-immobilized glucose oxidase and catalase. MSc thesis, Arizona State University, Tempe, AZ. Muehlbauer, M. J., Guilbeau, E. J. & Towe, B. C. (1989). Model for a thermoelectric enzyme glucose sensor. Anal. Chem., 61,77-83. Rich, S., Ianniello, R. M. & Jespersen, N. D. (1979). Development and application of a thermistor enzyme probe in the urea-urease system. Anal. Chem., 51, 204-6. Shichiri, M., Kawamori, R., Hakui, N., Asakawa, N., Yamasaki, Y. & Abe, H. (1984). The development of wearable-type artificial endocrine pancreas and its usefulness in glycaemic control of human diabetes mellitus. Biomed. Biochim. Actu, 5, 561-8.
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, 298-301.
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Tran-Minh, C. & Vallin, D. (1978). Enzyme-bound thermistor as an enthalpimetric sensor. Anal. Chem., 50, 187443. Tse, P. H. S. & Gough, D. A. (1987). Time-dependent inactivation of immobilized glucose oxidase and catalase. Biofech. Bioeng., XXIX, 705-13. Weaver, J. C., Cooney, C. L., Fulton, S. P., Schuler, P. & Tannenbaum, S. R. (1976). Experiments and calculations concerning a thermal enzyme probe. Biochim.
Biophys.
Acta, 452,285-91.
Thermoelectric Enzyme Sensor for Measuring Blood Glucose
Michael J. Muehlbauer,* Eric J. Guilbeau, Bruce C. Towe & Tedd A. Brandoi Arizona State University, Departmentof Chemical, Bio and MaterialsEngineering, Tempe, Arizona 85227,USA (Received 17February1989;accepted14 April 1989)
ABSTRACT A new calorimetric sensor has been developed which employs a thin-film thermopile in association with an immobilized enzyme. The thermopile detects the minute temperature rise that occurs when a specific chemical substrate is catalyzed by the enzyme. A prototype sensor is described which generates an equivalent proportional voltage response to glucose concentrations present in either buffer solution or blood. These sensors have remained useful for up to 18 days when operated intermittently for measuring glucose in buffer solutions, or for up to 4 days when operated continuously. When implanted inside cardiovascular shunts on anesthetized dogs, the sensors responded appropriately to changes in the blood glucose concentration. Key words: blood glucose, calorimetric, catalase, catheter, enzyme, glucose oxidase, implantation, thermoelectric, thermopile, stability.
INTRODUCTION The concept of a general calorimetric sensor for detecting biochemical reagents by virtue of their enzymatic enthalpy of reaction is attractive in terms of its versatility and universality. All previous sensors of this type *Presentaddress:Cytogam, Inc., 3498N San MarcosPlace, Suite 7, Chandler,Arizona 85224, USA. Biosensors & Bioelectronics 09%5663/89/$03-50
Ltd, England. Printed in Great Britain
0 1989 Elsevier Science Publishers
M. J. Muehlbauer,
2
E. J. Guilbeau, B. C. Towe, T. A. Brandon
have been fabricated with thermistors (Weaver et al., 1976; Tran-Minh &Vallin, 1978; Rich et al., 1979; Fulton & Cooney, 1980). While providing adequate sensitivity for the measurement of the extremely small temperature differences produced, thermistors have suffered in regards to their poor signal-to-noise ratio and problems associated with their necessity for an external voltage bias. An alternative calorimetric sensor has recently been advanced based on the thermoelectric principle (Guilbeau et al., 1988). With this device, enzyme is immobilized directly over one set of thermoelectric junctions on a thin-film thermopile. In response to a specific chemical substrate, the enzyme generates heat thereby establishing a temperature gradient across the thermopile. A voltage output results by virtue of the thermoelectric Seebeck effect whose magnitude is proportibnal to the concentration of substrate in the bulk solution. Initial experiments with thermoelectric enzyme sensors employed the enzyme glucose oxidase (GOD) coupled to the sensor surface to catalyze the reaction GOD
Glucose + O2
-+
Gluconic acid + H202 + 79 kJ/mol
Clearly, oxygen acts as a co-reactant and must remain in excess in order for the sensor response to be proportional to the concentration of glucose. Towards this purpose, a second enzyme, catalase, is co-immobilized to catalyze the breakdown of hydrogen peroxide created in the first reaction Catalase H202
-
H20 + l/2 O2 + 100 kJ/mol
The reaction regenerates oxygen and acts to extend the range of response of the sensor to glucose, and further provides an additional enthalpy of reaction which more than doubles the sensitivity. Previous results with the thermoelectric enzyme sensor indicated the dependence of its voltage output upon the amount of enzyme immobilized, and also upon the flow rate and temperature (Muehlbauer et al., 1989). Its response time, however, was extremely rapid (less than 6 s). Moreover, in comparison to enzyme-coupled thermistor devices, the signal-to-noise ratio and baseline stability were remarkably enhanced. Based on these characteristics, an investigation into the possibility of employing the thermoelectric enzyme sensor for measuring glucose in blood for applications in biomedicine or in clinical analysis was undertaken. The following reports the results of experiments for determining the response of the sensor in whole blood both in vitro and in a physiological environment within a cardiovascular shunt on an anesthetized dog. To
Thermoelectric enzyme sensor for measuring blood glucose
3
ascertain the applicability of the sensor for long-term glucose monitoring, we tested its stability in phosphate buffer solution both on a continuous basis and intermittently over a period of several days.
SENSOR FABRICATION The thermoelectric enzyme sensor was constructed in a probe form as illustrated in Fig. 1. Its fabrication has been described previously (Muehlbauer et al., 1989). Briefly, the thermopile consists of 50 couples of bismuth and antimony metal evaporated onto a Mylar (polyethylene terephthalate) insulating support. Individual thermopiles were cut and
Enzyme-coated thermopile
Shielded wire
Polyethylene
tube
EPOXY
\ Ground wire
Fig. 1. Thermoelectric
enzyme sensor.
rolled into cylindrical form for mounting at the tip of a 3-mm polyethylene catheter. The metals of the thermopile were positioned facing the interior of the cylinder and were bonded at their contact pads with shielded electrical wire. This wire extended through the catheter for external connection with a microvoltmeter. Polyurethane foam was sprayed through the interior of the catheter to retain the Mylar in a cylindrical shape. The seams of the Mylar were sealed with epoxy. A modification of the sensor for adaptation to blood measurements was the inclusion of additional electrical shielding over the thermopile. For this purpose, a thread of the wire-gauze shield covering the contact wire was positioned on the exterior surface of the catheter. The external surface of the Mylar insulation opposite to the thermopile was coated with a thin layer of silver conductive paint and connected with the wire shield. The shield was then appropriately grounded to reduce electrical interference effects. Enzyme was immobilized as previously described by first dipcoating the sensor in a solution of segmented polyurethane and then cross-linking a solution of GOD, catalase, and albumin with glutaraldehyde.
4
M. J. Muehlbauer, E. J. Guilbeau, B. C. Towe, T. A. Brandon
EXPERIMENTAL
METHODS
In-vitro experiments For the in-vitro experiments, sensors were tested in either phosphate buffer solution (pH 7.4) or in heparinized canine blood. The canine blood was exsanguinated from anesthetized dogs at the termination of in-vivo experiments and was reduced in glucose concentration by prior injection of insulin. A flow-loop apparatus was employed in which a peristaltic blood pump drew from a reservoir containing approximately 1 litre of either buffer solution or blood. The fluid passed through a loop of Tygon tubing inside which the sensor was inserted through one convergent leg of a Y-connector. Fluid flowed from the reservoir past the sensor in the same direction as the sensor tip was pointed. Prior to its return to the reservoir, the fluid was saturated with pure oxygen by passage through a bubble oxygenator. Voltage measurements from the sensor were obtained with a HewlettPackard null voltmeter and were measured relative to the zero baseline voltage of the amplifier. The voltage response to increasing concentrations of glucose was recorded following injections of a concentrated solution of glucose into the fluid reservoir. The concentration of glucose within the reservoir was measured after approximately 5 min of mixing by analyzing a sample on a Kodak Ektachem analyzer (Eastman Kodak Co., Rochester, New York, USA). Shunt implantation experiments For the purpose of testing the sensor in a physiological environment, a cardiovascular shunt was devised. The shunt consisted of a length of Tygon tubing for placement between the femoral artery and vein on opposite legs of anesthetized dogs. The sensor was again positioned in the flow by insertion through an in-line Y-connector. Separate Luer-Lok ports were also included in the shunt for sample withdrawal and for the insertion of transducers for temperature and pressure monitoring. Greyhound dogs (23-38 kg) were anesthetized intramuscularly with Ketamine Plus (11 mg ketamine, 2.2 mg xylazine, O-04 mg atropine per kg dog weight; Bristol Laboratories, Syracuse, New York, USA) followed by intravenous sodium pentobarbital. The shunt was placed following the surgical cutdown of the opposing femoral artery and vein, and the sensor was positioned in the direction of blood flow. The dog was respirated with pure oxygen with a Harvard positive volume respirator (Harvard Apparatus, South Natick, Massachusetts, USA) to maintain a blood oxygen
Thermoelectric enzyme sensor for measuring blood glucose
5
tension of roughly 400 mm Hg. Blood pH was maintained with injections of sodium bicarbonate. To ascertain the stability of the sensor in this physiological environment, the voltage output was recorded continuously and compared against glucose concentration measurements obtained from the Kodak analyzer on withdrawn blood samples. These samples were withdrawn every 10 or 15 min and were additionally analyzed for pH and blood gases with a Radiometer ABL2 analyzer (Radiometer Copenhagen, Copenhagen, Denmark). The response of the sensor to glucose was noted following the injection of a dose of insulin (Eli Lilly, Indianapolis, Indiana, USA; 1 U per kg dog weight). As a further test of stability, sensors were tested for their response to glucose in phosphate buffer solution both prior to and following the shunt experiments.
RESULTS Long-term stability tests
These tests were performed to ascertain the ability of the thermoelectric sensors to respond to glucose over a period of many days. Sensors were stored in buffer solutions at pH 7.4 and were periodically removed to test their response to glucose. Figure 2 displays the results for one particular sensor stored and tested in phosphate buffer at room temperature for a period of 18 days. Tests were performed with an oxygen saturated solution in which the concentration of glucose was progressively increased. The sensitivity to glucose for ADay1 *Days 2-4 *Days 6-l 1 n Day
4.54
13
VDays
16-18
4.
0
Fii.
50
100 150 GLUCOSE CONCENTRATION,
200 mg %
250
r 300
2. Long-term stability of thermoelectric enzyme glucose sensor stored and tested in phosphate buffer solution at room temperature.
6
M. J. Muehlbauer,
E. J. Guilbeau,
B. C. Towe, T. A. Brandon
this sensor decreased from an initial value of 15 nV/mg% on the first day of testing to a value of 6 nV/mg% on the eighteenth day. The range of linear response also declined from an initial value of 329 mg% to a final concentration of 140 mg%. The sensor failed to exhibit any response to glucose on the twentieth day of testing. In total, 6 individual sensors were tested for their long-term response to glucose. Three others were tested at room temperature-2 after storage at room temperature, a third with refrigerator storage. The remaining 2 sensors were stored at 37°C and tested at both room temperature and 37°C. These last sensors showed a larger response to glucose at the higher temperature, as expected based on the increased rate of glucose reaction. Typical characteristics for all sensors were a continual decline in both sensitivity and range of linear response. In general, the sensors responded to glucose for a longer period when stored at the lower temperatures. Sensors stored at room temperature remained useful for measuring glucose for 7,9, and 18 days; the sensor with refrigerator storage for 24 days; and those at 37°C for 5 and 6 days. The sensors typically failed by electrical failure, as characterized by a dramatic increase in signal noise which was believed to result from water permeation into the sensor core. The sensor represented in Fig. 2, however, did not exhibit this behavior and was believed to fail as a result of immobilized enzyme deactivation, as verified by an enzyme assay. Continuous
stability tests
Another set of stability tests were performed in which sensors were tested for their response to glucose during continuous operation. Here, each sensor was kept in a continuously flowing stream of buffer solution at room temperature and exposed to a constant concentration of glucose, typically 60 mg%. In all, 4 sensors were tested in this manner. In each case, the voltage output declined steadily over time from its initial value at first exposure to glucose. The normalized rate of decline, defined as the rate of voltage decline in V h-’ divided by the initial voltage response, was 0.0098 If:0.0030 h-‘. Three of the sensors appeared to fail by electrical failure and remained useful for periods of 30, 31, and 78 h. The fourth sensor remained useful for 97 h at which time its voltage response to glucose was reduced to zero, an apparent result of enzyme deactivation. In-vitro response to glucose in blood
Five different sensors were tested for their response to glucose in whole blood. The best results are displayed in Fig. 3. Here, the response to
Thermoelectric
enzyme sensor for measuring blood glucose
0 Phosphate buffer 0 Dog blood
solution
12+
c
,,,.,.,.,~,.,~r 0
Fig. 3. Comparison
50
100
200 250 150 OLUCOSE CONCENTRATION,
300 rng 96
350
400
450
of the response of the thermoelectric enzyme glucose sensor in blood and in phosphate buffer solution.
glucose in blood is compared to the response in phosphate buffer solution both before and after the blood measurements. The sensitivity to glucose in the buffer solution exhibited no change following the tests in blood, and remained stable at a value of 36 f 1 nV/mg%. The sensitivity in blood was 34 + 2 nV/mg%. The combined sensitivity to glucose was 35 + 1 nV/mg% and the range of linear response was 328 mg%. Other sensors tested for their response in whole blood often revealed a slightly different zero offset than they did when operated in buffer solution. In all cases, however, the sensitivity to glucose remained equivalent in the two solutions. The average sensitivity to glucose in blood was 42 k 10 nV/mg% compared to an average of 44 + 10 nV/mg% in phosphate buffer solution. Response to glucose in cardiovascular
shunts
Three different sensors were tested for their response to glucose when placed in cardiovascular shunts on anesthetized dogs. Figure 4 represents the response of one of these sensors which remained in the shunt for a period of 4 h. Here, the scale for the sensor output is converted into a glucose concentration scale with the aid of a linear regression calibration between the analyzer glucose measurements and the sensor voltage. Analyzer measurements of glucose performed on withdrawn blood samples are indicated in the figure as individual points. The calibrated glucose sensitivity was 47 nV/mg%. In Fig. 4, the response of the sensor following an injection of insulin into the dog is exhibited. The analyzer measure-
8
M. J. Muehlbauer, E. J. Guilbeau, B. C. Towe, T. A. Brandon INSULIN
IOmln
Fig. 4. Continuous
cardiovascular
output from a thermoelectric enzyme glucose sensor implanted in a shunt, indicating the response following injection of insulin. Vertical voltage scale is converted into concentration units.
Sensor
=
Glucose
I .Ol
Ektachem
Glucose
-
1.6
r2=
0.663
240. 220.
.
200. 180.
0
20
Fig. 5. Correlation
40
60
80 EKTACHEM
120 100 GLUCOSE, mg%
140
160
180
200
of glucose concentrations measured by the thermoelectric glucose sensor and by the Kodak Ektachem analyzer.
enzyme
ments indicate a decrease in glucose concentration from 120 to 30 mg% correlated with a decline in sensor voltage of approximately 4 pV. For all 3 shunt-implanted sensors combined, the average sensitivity to glucose measured in the blood stream was 40 + 7 nV/mg%. The sensitivity to glucose in phosphate buffer solution was comparable and dropped an average of 7% following the shunt implantation experiments. Figure 5 indicates the degree of correlation between the measurements of glucose concentration made with the 3 shunt-implanted sensors and those made on withdrawn blood samples with the Kodak analyzer. Because of the manner in which the sensors were calibrated, linear regression between the two measured concentrations of glucose produced a line with nearly unity slope. The value of ?, however, is significant in indicating the
Thermoelectric enzyme sensor for measuring blood glucose
9
uncertainty in the sensor measurement. As a result of this uncertainty, one particular value of glucose concentration as measured with the calibrated sensor was less than zero.
DISCUSSION The decline in sensor response to glucose over a period of many days as illustrated in Fig. 2 was typical for all sensors tested. In general, the initial linear sensitivity steadily declined as did the range of linear response, defined as the concentration of glucose at which the sensor output diverges from linearity. The former result was not surprising based on the rationale that enzyme deactivation reduced the rate of reaction and thereby the heat which the sensor detects. The second result was unexpected, however, based on simulations which suggested that a reduction in GOD activity should extend the range of linear response by allowing the reaction of glucose to be less diffusion limited (Muehlbauer et al., 1989). Apparently, the co-immobilized enzyme, catalase, was not present in a sufficient quantity to completely consume the hydrogen peroxide generated in the GOD-catalyzed reaction. Hydrogen peroxide causes deactivation of both enzymes (Buchholz & Godelmann, 1978; Tse & Gough, 1987) and is preferably eliminated by providing sufficient excess of catalase. However, as separate experiments with co-immobilized membranes of GOD and catalase illustrate, even an 80-fold excess in catalase actvity was not sufficient to achieve maximum enzyme stability (Ireland, 1988). The indication is that some hydrogen peroxide remains unconsumed. Without complete consumption of hydrogen peroxide, a stoichiometric ratio of oxygen to glucose consumption of more than 1:2 results. Furthermore, since catalase deactivates faster than GOD, the ratio is expected to increase over time. Consequently, the range of linear response for the sensor is reduced as oxygen becomes the limiting reactant at increasingly lower concentrations of glucose. Likewise, the sensitivity is further reduced as less heat is generated in the coupled catalase reaction. These observations are consistent with the results of the long-term stability tests. As with the long-term stability tests, sensors tested on a continuous basis showed a steady although more rapid decline in sensitivity. For both tests and in most cases, the sensors appeared to fail as a result of water permeation into the thin-film element of the device. This occurred despite the excellent hydrophobicity of the Mylar and the presence of the epoxy sealant. Upon failure, sensors often revealed an abrupt decrease in
10
M. J. Muehlbauer,
E. J. Guilbeau,
B. C. Towe, T. A. Brandon
signal-to-noise ratio and an inability to respond to glucose even though a subsequent enzyme assay revealed the presence of enzyme activity. These sensors often exhibited an increase in resistance, although not always to an infinite value. Other sensors appeared to fail by virtue of enzyme deactivation as indicated by a loss of assayed enzyme activity and a maintenance of good electrical characteristics. Upon initiation of the experiments for testing the sensor performance in blood both in vitro and in the shunt, a noise component in the signal was immediately discovered which was not present for the tests in buffer solution. The additional noise was partially attributed to thermal interference, but was also an apparent interaction of the sensor with the blood. It was postulated that electrical interference from streaming currents caused by the adherence of charged proteins was the cause. The addition of an electrical shield opposite the thin-film thermopile improved the signal output from the sensor for all experiments in which the sensor was immersed in blood. As indicated in the in-vitro blood experiments, the sensor response to glucose generally exhibited the same sensitivity whether the sensor was operated in blood or in buffer solution. This equivalence in sensitivity, as Fig. 3 typifies, was found when the concentration of glucose in blood was measured as the plasma value. The analytical measurements for glucose in blood in this figure were measured by spinning the samples in a centrifuge prior to determination on the Kodak Ektachem analyzer. The indication, therefore, is that the sensor appeared to respond to the plasma concentration of glucose despite its operation in whole blood. This is consistent with the consideration that the sensor operates in a steady-state mode and draws its glucose for measurement from the plasma. Such a relatively small amount is consumed that glucose from the blood cells is not extracted. The sensor response in the cardiovascular shunt was acted upon by considerable interference. Among the probable sources were thermal gradients, piezoelectric currents caused by the blood pressure, and bioelectric currents. The comparison between the analytical and the sensor measurements of glucose, as Fig. 4 illustrates, was obscured by noise, sensor voltage offset, and a minimal sensitivity to glucose for the thermoelectric sensor. Despite these complications, however, a discernible response to glucose was detected. The response following a decline in glucose concentration as initiated by insulin administration was consistent with expectation. In addition, although the sensor performed somewhat erratically in the shunt environment, the exposure did not appear to produce any serious deleterious effects. The sensor performed equally well both before and after implantation when operated in buffer solution. The reduction in sensitiv-
Thermoelectric enzyme sensor for measuring blood glucose
11
ity following implantation was no greater than that found for sensors operated continuously in buffer solutions. The thermoelectric enzyme glucose sensor represents the first calorimetric device to be implanted and operated in a continuous mode in the blood stream. Realistically, the sensor cannot compete with polarographic or optical enzyme sensors for similar applications. Disadvantages include its present size and reduced sensitivity. These experiments indicate, however, that anticipated problems associated with thermal and electrical interference and stability can be overcome. In truth, the signal-to-noise ratio of the thermoelectric sensor is comparable to that of other implanted glucose sensors (Shichiri et al., 1984, 1986). Its improved response time and adaptability to other enzyme-substrate systems therefore make it attractive for analytical purposes, if not for biomedical implantation. ACKNOWLEDGEMENTS This research was supported by grants from the Whitaker Foundation the Arizona Disease Control Research Commission.
and
REFERENCES Buchholz, K. & Godelmann, B. (1978). Macrokinetics and operational stability of immobilized glucose oxidase and catalase. Biotech. Bioeng., XX, 1201-20. Fulton, S. P. & Cooney, C. L. (1980). Thermal enzyme probe with differential temperature measurements in a laminar flow-through cell. Anal. Chem., 52, 505-8.
Guilbeau, E. J., Towe, B. C. & Muehlbauer, M. J. (1988). A potentially implantable thermoelectric sensor for measurement of glucose. Trans. Am. Sot. Artif. Intern. Organs, XxX111, 329-35. Ireland, T. W. (1988). In vitro and in vivo stability of co-immobilized glucose oxidase and catalase. MSc thesis, Arizona State University, Tempe, AZ. Muehlbauer, M. J., Guilbeau, E. J. & Towe, B. C. (1989). Model for a thermoelectric enzyme glucose sensor. Anal. Chem., 61,77-83. Rich, S., Ianniello, R. M. & Jespersen, N. D. (1979). Development and application of a thermistor enzyme probe in the urea-urease system. Anal. Chem., 51, 204-6. Shichiri, M., Kawamori, R., Hakui, N., Asakawa, N., Yamasaki, Y. & Abe, H. (1984). The development of wearable-type artificial endocrine pancreas and its usefulness in glycaemic control of human diabetes mellitus. Biomed. Biochim. Actu, 5, 561-8.
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, 298-301.
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Tran-Minh, C. & Vallin, D. (1978). Enzyme-bound thermistor as an enthalpimetric sensor. Anal. Chem., 50, 187443. Tse, P. H. S. & Gough, D. A. (1987). Time-dependent inactivation of immobilized glucose oxidase and catalase. Biofech. Bioeng., XXIX, 705-13. Weaver, J. C., Cooney, C. L., Fulton, S. P., Schuler, P. & Tannenbaum, S. R. (1976). Experiments and calculations concerning a thermal enzyme probe. Biochim.
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