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Reproducible electrodeposition of biomolecules for the fabrication of miniature electroenzymatic biosensors
Sensors and Actuators
B, 5 (1991) 85-89
Reproducible electrodeposition of biomolecules for the fabrication of miniature electroenzymatic biosensors Kirk W. Johnson Medical Device and Diagnostics Indianapolis, IN 4628.5 (U.S.A.)
Division, Eli Lilly and Company,
Lilly Corporate
Center, Drop Code 081 I,
Abstract A technique has been developed for reproducibly depositing an enzyme or other biomolecule onto the surface of a conductor. Specifically, glucose oxidase has been electrodeposited onto the surface of an electrode previously electroplated with platinum black. This deposition takes place without the assistance of an electropolymerizable matrix material such as polypyrrole. The deposition process does not require a time-consuming alignment step, characteristic of some deposition techniques utilizing photoimageable matrix materials; yet it results in the enzyme being precisely deposited only on the working electrode. The thickness of the resulting enzyme layer can be varied from Angstroms to microns by varying the deposition parameters. With proper fixturing, potentially hundreds of sensors can be. simultaneously coated with enzyme.
Introduction
The immobilization of biomolecules such as enzymes and antibodies is a necessary procedure for many analytical applications. Immobilization renders the molecules waterinsoluble by attaching them to a physical support material or themselves. This procedure then allows the molecules to be reused many times without a change in concentration or activity. Immobilization also allows the biomolecules to be retained in a specific region of the apparatus or sensor. A variety of techniques has been utilized for the immobilization of enzymes for use in biosensors [l-8]. Unfortunately, very few of these techniques are adaptable to the miniature electrochemical biosensors currently being fabricated using semiconductor techniques. These semiconductor techniques allow a sensor to be made with a surface area of 1.0 pm*. For the fabrication of enzyme sensors, it is desirable to deposit accurately and immobilize the enzyme only on the surface of the working electrode, which is used to monitor a product of the enzymatic reaction. Enzyme deposited on other areas of the sensor would be reacting to produce a
09254005/91/%3.50
product that could not be detected, which is useless and should be minimized due to the expense of the enzyme. For the development of these techniques, it is important to appreciate that the enzymes are very fragile molecules in that their activity is easily degraded upon exposure to most organic chemicals, temperatures above approximately 50 “C, and extreme pH (high or low). These considerations limit the arsenal of available techniques for enzyme deposition. Kimura et al. [9] developed a method for enzyme deposition using lift-off techniques with photoresist. Heineman and others have evaluated the potential of gamma irradiation as a means of immobilizing an enzyme in a polymer matrix on an electrode surface [lo]. This method can also be used with chemicals that polymerize upon exposure to ultraviolet light. Shinohara et al. have synthesized enzyme-containing membrane films by electrochemical polymerization of pyrrole or aniline in the presence of enzyme [ 111. Platinized platinum, sometimes referred to as ‘platinum black’, was employed as the supporting matrix for enzyme deposition by Ikariyama et al. [ 121. The enzyme is incorporated into these pores by immersing the platinum-blacked
@ 1991 -
Elsevier Sequoia, Lausanne
86
electrode in a phosphate buffered solution (pH = 7.0) containing the enzyme. The enzyme molecules adsorb to the platinum black surface and can be further anchored to the sensor by covalent crosslinking with glutaraldehyde. This group also modified their procedure by adding the enzyme to the hexachloroplatinate solution and simultaneously electrodepositing platinum black and glucose oxidase [ 131. A technique has been developed for reproducibly depositing an enzyme or other biomolecule onto the surface of a conductor. Specifically, glucose oxidase has been electrodeposited onto the surface of an electrode previously electroplated with platinum black. This deposition takes place in an aqueous solution (pH = 7.4) without the assistance of an electropolymerizable matrix material such as polypyrrole. The deposition process does not require the time-consuming alignment step that is characteristic of some deposition techniques utilizing photoimageable matrix materials; yet it results in the enzyme being precisely deposited only on the working electrode. The thickness of the resulting enzyme layer can be varied from Angstroms to microns by varying the deposition parameters. With proper fixturing, potentially hundreds of sensors can be simultaneously coated with enzyme.
Theory
The electrophoretic migration of charged molecules, particularly proteins, in an electric field is well known and documented in textbooks [ 141. This migration is based on the isoelectric point of the molecule, which is the pH at which the positive and negative charges on the molecule are exactly balanced and no migration takes place. At a pH value greater than the isoelectric point of the protein molecule, the molecule possesses a net negative charge, and vice versa. Glucose oxidase has an isoelectric point of 4.3, which results in it exhibiting a net negative charge in a phosphate-buffered saline (PBS) solution
(pH = 7.4). This pH is very hospitable for the glucose oxidase and denaturing is not observed. The negatively charged glucose oxidase molecule is preferentially attracted to an oppositely charged electrode, followed by covalent crosslinking to anchor the glucose oxidase molecules onto the surface of the electrode. Multiple proteins can be codeposited if their isoelectric points all lie on the same side (i.e., higher or lower) of the solution pH.
Experimental
The amperometric sensor onto which the glucose oxidase is deposited is fabricated using thin-/thick-film processing techniques [ 151. This process results in a three-electrode sensor with a 0.1 mm* platinum-black working electrode. The sensor also contains a platinum-blacked counter electrode and an Ag/AgCl reference electrode. A 5% by weight solution of glucose oxidase is prepared utilizing a phosphate-buffered saline solution (pH 7.4) as the solvent. While dissolving the glucose oxidase by stirring, it is important to prevent foaming of the solution which could denature the glucose oxidase. The electrodeposition cell is filled with this solution and the electrode onto which the enzyme is to be deposited, along with a counter electrode, are submerged inside the cell. Electrical connections are made to a galvanostat and current is applied such that a constant density of 5 mA/cm* results at the electrode surface. The current flow is oriented in a direction to result in a positive charge on the electrode surface to attract the negatively charged glucose oxidase molecules. The current is applied for two minutes. The actual voltage at the working electrode varies with the impedance. The electrode is removed from the enzyme solution and submerged in deionized distilled water with no agitation for five seconds to remove residual glucose oxidase. Following this, the electrode is submerged in a solution containing 2.5% by volume glutaraldehyde in phosphate-buffered
87
saline (pH 7.4) for 30 minutes to crosslink the glucose m,+cules covalently together and form a water-ifisoluble layer. The glutaraldehyde crosslinking prevents the glucose oxidase from going back into solution. The electrode is again submerged in deionized distilled water for five seconds and allowed to dry in air for 30 minutes. The enzymatic component of the biosensor is functional at this point. Another important aspect of this method is the capability to electrodeposit simultaneously two or more different biomolecules providing their isoelectric points are all either higher or lower than the fiH of the deposition solution. A solution is prepared containing 5% by weight glucose oxidase and 5% by weight albumin (bovine) in a phosphatebuffered saline solution (pH 7.4). The glucose oxidase and albumin have isoelectric point values of 4.3 and 4.7 respectively, resulting in both molecules possessing a negative charge in the pH 7.4 solution. A current density of 5.0 mA/cm2 is applied to the electrode for two minutes while it is submerged in the glucose oxidase/albumin solution. The molecules are crosslinked in the glutaraldehyde solution as described above, rinsed and dried. Additives such as non-ionic surfactants (Union Carbide Corp.) and anti-foaming agents (Dow Corning) have been added to
ADSORPTION 500
Fig. 1. Comparison of current output for three sensors in a 100 mg/dl glucose solution in vitro. The sensors are identical in electrode area and method of fabrication, except for the enzyme deposition. One of the sensors is coated using the Ikariyama adsorption technique [12] and the other two are coated using the method described here utilizing two different current values. The sensors are not covered with an outer membrane.
the solution with no loss of enzyme activity in the resulting deposited layer. The in vitro performance of the sensors is evaluated by submersion in phosphatebuffered saline solutions (Sigma Chemical Company), pH 7.4, with the temperature maintained at 37 “C. The glucose concentration is adjusted by spiking with a concentrated (10 000 mg/dl) dextrose solution that has been at room temperature for 24 hours prior to use to assure complete mutarotation of the glucose. A potential of +0.6 V (versus Ag/AgCl) is applied to the sensors using a potentiostat fabricated in our laboratories.
Results and discussion The conditions outlined above normally result in a glucose oxidase layer approximately 1.5 pm thick when dry as measured with a profilometer (Dektak). The amount of enzyme deposited can be changed by varying the current density during the electrodeposition step. Figure 1 illustrates the resulting current from electrodes with identical surface area coated using the Ikariyama technique [ 121 and the method described here utilizing two different current levels. For this test, the sensors did not have an outer membrane on top of the enzyme layer, so the current from the sensor is a direct indication of the amount of glucose oxidase immobilized on the surface of the electrode. From these data, it is evident that the application of current to the electrode enhances the attraction and subsequent attachment of glucose oxidase to the surface of the working electrode compared to simple adsorption. In addition, the amount of glucose oxidase deposited on the electrode is proportional to the current density. Care must be taken when selecting the current density so that the corresponding voltage does not climb to a level where deleterious gas bubbles would be generated, resulting in bubbles entrapped in the biomolecule layer. The codeposition of glucose oxidase and albumin resulted in a layer approximately 5 pm thick. This combination of biomole-
88
0
10
20
-
433Dl
-
433D?.
-
433DS
-
43x86
30
40
TIME (HRS) Fig. Current vs. a group of four sensors in a mg/dl solution in vitro. The working electrode has glucose oxidase only electrodeposited on its surface, followed by crosslinking in glutaraldehyde. The sensors are covered with an outer membrane.
Fig. 4. SEM photograph (400 x , reduced in reproduction 75%) of a portion of the working electrode that has glucose oxidase and albumin electrodeposited and crosslinked onto the exposed surface. The visible metal layer surrounding the layer of glucose oxidase and albumin is the conductor underlying the insulation layer. This sensor is not covered with an outer membrane. 01 0
20
40
60
80
TIME (HRS) Fig. 3. Current output vs. time for a group of four sensors in a 100 mg/dl glucose solution in vitro. The working electrode had glucose oxidase and albumin codeposited on its surface, followed by crosslinking in glutaraldehyde. The sensors are covered with an outer membrane.
cules results in sensors which had a more stable long-term output in a 100 mg/dl glucose solution than glucose oxidase alone, as illustrated in Figs. 2 and 3. The sensors having glucose oxidase and albumin immobilized together do not exhibit the downward drift over time seen with the sensors comprising glucose oxidase only. This long-term stability over a 72 hour period is absolutely necessary for an implantable biosensor. This enhanced stability may be due to the formation of a protein layer that was more tightly crosslinked upon exposure to glutaraldehyde and did not allow glucose oxidase to leach
out. In addition, the albumin may have provided an environment that was more hospitable to the glucose oxidase in terms of hydrophobicity, localized pH, etc. The deposition of the enzyme is very sitespecific, as shown in the scanning electron micrograph, Fig. 4. This micrograph illustrates the enzyme layer deposited on top of the electrode only in the region where the top insulation layer was removed. This deposition of the enzyme only on the working electrode minimizes the amount of glucose oxidase wasted and thus minimizes the production costs. Due to the electrical nature of this process, large numbers of sensors may be coated with biomolecules in a simultaneous, parallel manner. In addition, the electrodeposition process allows the coating of conductors with uneven topographies or geometrical surfaces, such as the inside of metal tubes, for use in flowthrough processes or HPLC columns.
89
Following coating of the sensors with a proprietary outer membrane material, the in vitro characteristics of the sensors were determined and have been published elsewhere
[151. Conclusions
In conclusion, a method has been developed for reproducibly electrodepositing an enzyme or other biomolecule onto the surface of conductors, regardless of the conductors’ size and topography. This deposition takes place from an aqueous solution whose pH can be matched to the hospitable range for the biomolecule. The electrodeposition process allows potentially hundreds of sensors to be coated simultaneously, yet very precisely and with controllable thicknesses.
Acknowledgements
6 U. Fischer and P. Abel, A membrane
combination for implantable glucose sensor measurements in undiluted biological fluids, Trans. Am. Sot. Artif.
Intern. Organs, 28 (1982) 245-248. 7 M. Shichiri, R. Kawamori, Y. Goriya,
Y. Yamasaki, M. Nomura, N. Hakui and H. Abe, Glycaemic control in pancreatectomized dogs with a wearable artifical endocrine pancreas, Diabetologia,
24 (1983) 179-184. 8 L. Gorton and G. Jonsson, A glucose sensor based
on the adsorption of glucose on a palladium/gold modified carbon electrode, J. Mol. Catal., 38 (1986) 157-159. 9 J. Kimura,
N. Ito, K. Toshihide, M. Kikuchi, T. Arai, N. Negishi and Y. Tomita, A novel blood glucose monitoring method: an ISFET biosensor applied to transcutaneous effusion fluid, J. Elec-
trochem. Sot., 136 (6) (1989) 17441747. IO W. R. Heineman, K. Hajizadeh, R. S. Tieman,
K. L. Rauen and L. A. Coury, Jr., Sensors based on polymer networks formed by gamma radiation crosslinking, Proc. Third Int. Meeting Chem. Sensors, Cleveland, OH, U.S.A., Sept. 24-26, 1990, pp. 369-372.
11 H. Shinohara, T. Chiba and M. Aizawa, Enzyme microsensor for glucose with an eiectrochemically synthesized enzyme-polyaniline film, Sensors and Actuators,
The author wishes to thank Doug Allen and Nan Cox for helpful suggestions and discussions regarding surfactants and antifoaming agents. In addition, the fabrication of sensors by Charles Andrew and potentiostats by William McMahan is greatly appreciated
References S. J. Updike and G. P. Hicks, The enzyme electrode, Nature, 214 (1967) 986-988.
L. C. Clark, Jr., R. B. Spokane, R. Sudan and T. L. Stroup, Long-lived implanted silastic drum glucose sensors, Trans. Am. Sot. Artif. Intern. Organs, 33 (1987) 323-328.
D. A. Gough, J. Y. Luciasana and P. H. S. Tse, Two-dimensional enzyme electrode sensor for glucose, Anal. Chem., 57 (1985) 2351-2357. J. L. Romette, B. Froment and D. Thomas, Glucase-oxidase electrode. Measurements of glucose in samples exhibiting high variability in oxygen content, Clin. Chirn. Acta, 9.5 (1979) 249-253. D. J. Claremont, I. E. Sambrook, C. Penton and J. C. Pickup, Subcutaneous implantation of a ferrocene-mediated glucose sensor in pigs, Diabetologia, 29 (1986) 817-821.
I3 (1988) 79-86.
12 Y. Ikariyama, S. Yamauchi, T. Yukiashi and H. Ushioda, Micro-enzyme electrode prepared on platinized platinum, Anal. Lett., 20 (9) (1987) 14071416.
13 Y. Ikariyama, S. Yamauchi, T. Yukiashi and H. Ushioda, Electrochemical fabrication of amperometric microenzyme sensor, J. Electrochem. Sot. I36 (3) ( 1989) 702-706.
14 R. T. Morrison and R. N. Boyd, Organic Chemistry, Allyn and Bacon, Inc., Boston, MA, 3rd edn., 1973, p. 1137. I5 J. J. Mastrototaro, K. W. Johnson, R. J. Morff, D. Lipson, C. C. Andrew and D. J. Allen, An electroenzymatic glucose sensor fabricated on a flexible substrate, Sensors and Actuators B, 5 (1991) 139-144.
Biography Kirk Johnson obtained B.S. and M.A. degrees in chemistry from the University of South Dakota, Vermillion in 1984 and 1986, respectively. He is a chemist in the Medical Device and Diagnostics Division of Eli Lilly and Company. His scientific interests include biosensors, electroanalytical chemistry and electrogenerated chemiluminescence reactions.
B, 5 (1991) 85-89
Reproducible electrodeposition of biomolecules for the fabrication of miniature electroenzymatic biosensors Kirk W. Johnson Medical Device and Diagnostics Indianapolis, IN 4628.5 (U.S.A.)
Division, Eli Lilly and Company,
Lilly Corporate
Center, Drop Code 081 I,
Abstract A technique has been developed for reproducibly depositing an enzyme or other biomolecule onto the surface of a conductor. Specifically, glucose oxidase has been electrodeposited onto the surface of an electrode previously electroplated with platinum black. This deposition takes place without the assistance of an electropolymerizable matrix material such as polypyrrole. The deposition process does not require a time-consuming alignment step, characteristic of some deposition techniques utilizing photoimageable matrix materials; yet it results in the enzyme being precisely deposited only on the working electrode. The thickness of the resulting enzyme layer can be varied from Angstroms to microns by varying the deposition parameters. With proper fixturing, potentially hundreds of sensors can be. simultaneously coated with enzyme.
Introduction
The immobilization of biomolecules such as enzymes and antibodies is a necessary procedure for many analytical applications. Immobilization renders the molecules waterinsoluble by attaching them to a physical support material or themselves. This procedure then allows the molecules to be reused many times without a change in concentration or activity. Immobilization also allows the biomolecules to be retained in a specific region of the apparatus or sensor. A variety of techniques has been utilized for the immobilization of enzymes for use in biosensors [l-8]. Unfortunately, very few of these techniques are adaptable to the miniature electrochemical biosensors currently being fabricated using semiconductor techniques. These semiconductor techniques allow a sensor to be made with a surface area of 1.0 pm*. For the fabrication of enzyme sensors, it is desirable to deposit accurately and immobilize the enzyme only on the surface of the working electrode, which is used to monitor a product of the enzymatic reaction. Enzyme deposited on other areas of the sensor would be reacting to produce a
09254005/91/%3.50
product that could not be detected, which is useless and should be minimized due to the expense of the enzyme. For the development of these techniques, it is important to appreciate that the enzymes are very fragile molecules in that their activity is easily degraded upon exposure to most organic chemicals, temperatures above approximately 50 “C, and extreme pH (high or low). These considerations limit the arsenal of available techniques for enzyme deposition. Kimura et al. [9] developed a method for enzyme deposition using lift-off techniques with photoresist. Heineman and others have evaluated the potential of gamma irradiation as a means of immobilizing an enzyme in a polymer matrix on an electrode surface [lo]. This method can also be used with chemicals that polymerize upon exposure to ultraviolet light. Shinohara et al. have synthesized enzyme-containing membrane films by electrochemical polymerization of pyrrole or aniline in the presence of enzyme [ 111. Platinized platinum, sometimes referred to as ‘platinum black’, was employed as the supporting matrix for enzyme deposition by Ikariyama et al. [ 121. The enzyme is incorporated into these pores by immersing the platinum-blacked
@ 1991 -
Elsevier Sequoia, Lausanne
86
electrode in a phosphate buffered solution (pH = 7.0) containing the enzyme. The enzyme molecules adsorb to the platinum black surface and can be further anchored to the sensor by covalent crosslinking with glutaraldehyde. This group also modified their procedure by adding the enzyme to the hexachloroplatinate solution and simultaneously electrodepositing platinum black and glucose oxidase [ 131. A technique has been developed for reproducibly depositing an enzyme or other biomolecule onto the surface of a conductor. Specifically, glucose oxidase has been electrodeposited onto the surface of an electrode previously electroplated with platinum black. This deposition takes place in an aqueous solution (pH = 7.4) without the assistance of an electropolymerizable matrix material such as polypyrrole. The deposition process does not require the time-consuming alignment step that is characteristic of some deposition techniques utilizing photoimageable matrix materials; yet it results in the enzyme being precisely deposited only on the working electrode. The thickness of the resulting enzyme layer can be varied from Angstroms to microns by varying the deposition parameters. With proper fixturing, potentially hundreds of sensors can be simultaneously coated with enzyme.
Theory
The electrophoretic migration of charged molecules, particularly proteins, in an electric field is well known and documented in textbooks [ 141. This migration is based on the isoelectric point of the molecule, which is the pH at which the positive and negative charges on the molecule are exactly balanced and no migration takes place. At a pH value greater than the isoelectric point of the protein molecule, the molecule possesses a net negative charge, and vice versa. Glucose oxidase has an isoelectric point of 4.3, which results in it exhibiting a net negative charge in a phosphate-buffered saline (PBS) solution
(pH = 7.4). This pH is very hospitable for the glucose oxidase and denaturing is not observed. The negatively charged glucose oxidase molecule is preferentially attracted to an oppositely charged electrode, followed by covalent crosslinking to anchor the glucose oxidase molecules onto the surface of the electrode. Multiple proteins can be codeposited if their isoelectric points all lie on the same side (i.e., higher or lower) of the solution pH.
Experimental
The amperometric sensor onto which the glucose oxidase is deposited is fabricated using thin-/thick-film processing techniques [ 151. This process results in a three-electrode sensor with a 0.1 mm* platinum-black working electrode. The sensor also contains a platinum-blacked counter electrode and an Ag/AgCl reference electrode. A 5% by weight solution of glucose oxidase is prepared utilizing a phosphate-buffered saline solution (pH 7.4) as the solvent. While dissolving the glucose oxidase by stirring, it is important to prevent foaming of the solution which could denature the glucose oxidase. The electrodeposition cell is filled with this solution and the electrode onto which the enzyme is to be deposited, along with a counter electrode, are submerged inside the cell. Electrical connections are made to a galvanostat and current is applied such that a constant density of 5 mA/cm* results at the electrode surface. The current flow is oriented in a direction to result in a positive charge on the electrode surface to attract the negatively charged glucose oxidase molecules. The current is applied for two minutes. The actual voltage at the working electrode varies with the impedance. The electrode is removed from the enzyme solution and submerged in deionized distilled water with no agitation for five seconds to remove residual glucose oxidase. Following this, the electrode is submerged in a solution containing 2.5% by volume glutaraldehyde in phosphate-buffered
87
saline (pH 7.4) for 30 minutes to crosslink the glucose m,+cules covalently together and form a water-ifisoluble layer. The glutaraldehyde crosslinking prevents the glucose oxidase from going back into solution. The electrode is again submerged in deionized distilled water for five seconds and allowed to dry in air for 30 minutes. The enzymatic component of the biosensor is functional at this point. Another important aspect of this method is the capability to electrodeposit simultaneously two or more different biomolecules providing their isoelectric points are all either higher or lower than the fiH of the deposition solution. A solution is prepared containing 5% by weight glucose oxidase and 5% by weight albumin (bovine) in a phosphatebuffered saline solution (pH 7.4). The glucose oxidase and albumin have isoelectric point values of 4.3 and 4.7 respectively, resulting in both molecules possessing a negative charge in the pH 7.4 solution. A current density of 5.0 mA/cm2 is applied to the electrode for two minutes while it is submerged in the glucose oxidase/albumin solution. The molecules are crosslinked in the glutaraldehyde solution as described above, rinsed and dried. Additives such as non-ionic surfactants (Union Carbide Corp.) and anti-foaming agents (Dow Corning) have been added to
ADSORPTION 500
Fig. 1. Comparison of current output for three sensors in a 100 mg/dl glucose solution in vitro. The sensors are identical in electrode area and method of fabrication, except for the enzyme deposition. One of the sensors is coated using the Ikariyama adsorption technique [12] and the other two are coated using the method described here utilizing two different current values. The sensors are not covered with an outer membrane.
the solution with no loss of enzyme activity in the resulting deposited layer. The in vitro performance of the sensors is evaluated by submersion in phosphatebuffered saline solutions (Sigma Chemical Company), pH 7.4, with the temperature maintained at 37 “C. The glucose concentration is adjusted by spiking with a concentrated (10 000 mg/dl) dextrose solution that has been at room temperature for 24 hours prior to use to assure complete mutarotation of the glucose. A potential of +0.6 V (versus Ag/AgCl) is applied to the sensors using a potentiostat fabricated in our laboratories.
Results and discussion The conditions outlined above normally result in a glucose oxidase layer approximately 1.5 pm thick when dry as measured with a profilometer (Dektak). The amount of enzyme deposited can be changed by varying the current density during the electrodeposition step. Figure 1 illustrates the resulting current from electrodes with identical surface area coated using the Ikariyama technique [ 121 and the method described here utilizing two different current levels. For this test, the sensors did not have an outer membrane on top of the enzyme layer, so the current from the sensor is a direct indication of the amount of glucose oxidase immobilized on the surface of the electrode. From these data, it is evident that the application of current to the electrode enhances the attraction and subsequent attachment of glucose oxidase to the surface of the working electrode compared to simple adsorption. In addition, the amount of glucose oxidase deposited on the electrode is proportional to the current density. Care must be taken when selecting the current density so that the corresponding voltage does not climb to a level where deleterious gas bubbles would be generated, resulting in bubbles entrapped in the biomolecule layer. The codeposition of glucose oxidase and albumin resulted in a layer approximately 5 pm thick. This combination of biomole-
88
0
10
20
-
433Dl
-
433D?.
-
433DS
-
43x86
30
40
TIME (HRS) Fig. Current vs. a group of four sensors in a mg/dl solution in vitro. The working electrode has glucose oxidase only electrodeposited on its surface, followed by crosslinking in glutaraldehyde. The sensors are covered with an outer membrane.
Fig. 4. SEM photograph (400 x , reduced in reproduction 75%) of a portion of the working electrode that has glucose oxidase and albumin electrodeposited and crosslinked onto the exposed surface. The visible metal layer surrounding the layer of glucose oxidase and albumin is the conductor underlying the insulation layer. This sensor is not covered with an outer membrane. 01 0
20
40
60
80
TIME (HRS) Fig. 3. Current output vs. time for a group of four sensors in a 100 mg/dl glucose solution in vitro. The working electrode had glucose oxidase and albumin codeposited on its surface, followed by crosslinking in glutaraldehyde. The sensors are covered with an outer membrane.
cules results in sensors which had a more stable long-term output in a 100 mg/dl glucose solution than glucose oxidase alone, as illustrated in Figs. 2 and 3. The sensors having glucose oxidase and albumin immobilized together do not exhibit the downward drift over time seen with the sensors comprising glucose oxidase only. This long-term stability over a 72 hour period is absolutely necessary for an implantable biosensor. This enhanced stability may be due to the formation of a protein layer that was more tightly crosslinked upon exposure to glutaraldehyde and did not allow glucose oxidase to leach
out. In addition, the albumin may have provided an environment that was more hospitable to the glucose oxidase in terms of hydrophobicity, localized pH, etc. The deposition of the enzyme is very sitespecific, as shown in the scanning electron micrograph, Fig. 4. This micrograph illustrates the enzyme layer deposited on top of the electrode only in the region where the top insulation layer was removed. This deposition of the enzyme only on the working electrode minimizes the amount of glucose oxidase wasted and thus minimizes the production costs. Due to the electrical nature of this process, large numbers of sensors may be coated with biomolecules in a simultaneous, parallel manner. In addition, the electrodeposition process allows the coating of conductors with uneven topographies or geometrical surfaces, such as the inside of metal tubes, for use in flowthrough processes or HPLC columns.
89
Following coating of the sensors with a proprietary outer membrane material, the in vitro characteristics of the sensors were determined and have been published elsewhere
[151. Conclusions
In conclusion, a method has been developed for reproducibly electrodepositing an enzyme or other biomolecule onto the surface of conductors, regardless of the conductors’ size and topography. This deposition takes place from an aqueous solution whose pH can be matched to the hospitable range for the biomolecule. The electrodeposition process allows potentially hundreds of sensors to be coated simultaneously, yet very precisely and with controllable thicknesses.
Acknowledgements
6 U. Fischer and P. Abel, A membrane
combination for implantable glucose sensor measurements in undiluted biological fluids, Trans. Am. Sot. Artif.
Intern. Organs, 28 (1982) 245-248. 7 M. Shichiri, R. Kawamori, Y. Goriya,
Y. Yamasaki, M. Nomura, N. Hakui and H. Abe, Glycaemic control in pancreatectomized dogs with a wearable artifical endocrine pancreas, Diabetologia,
24 (1983) 179-184. 8 L. Gorton and G. Jonsson, A glucose sensor based
on the adsorption of glucose on a palladium/gold modified carbon electrode, J. Mol. Catal., 38 (1986) 157-159. 9 J. Kimura,
N. Ito, K. Toshihide, M. Kikuchi, T. Arai, N. Negishi and Y. Tomita, A novel blood glucose monitoring method: an ISFET biosensor applied to transcutaneous effusion fluid, J. Elec-
trochem. Sot., 136 (6) (1989) 17441747. IO W. R. Heineman, K. Hajizadeh, R. S. Tieman,
K. L. Rauen and L. A. Coury, Jr., Sensors based on polymer networks formed by gamma radiation crosslinking, Proc. Third Int. Meeting Chem. Sensors, Cleveland, OH, U.S.A., Sept. 24-26, 1990, pp. 369-372.
11 H. Shinohara, T. Chiba and M. Aizawa, Enzyme microsensor for glucose with an eiectrochemically synthesized enzyme-polyaniline film, Sensors and Actuators,
The author wishes to thank Doug Allen and Nan Cox for helpful suggestions and discussions regarding surfactants and antifoaming agents. In addition, the fabrication of sensors by Charles Andrew and potentiostats by William McMahan is greatly appreciated
References S. J. Updike and G. P. Hicks, The enzyme electrode, Nature, 214 (1967) 986-988.
L. C. Clark, Jr., R. B. Spokane, R. Sudan and T. L. Stroup, Long-lived implanted silastic drum glucose sensors, Trans. Am. Sot. Artif. Intern. Organs, 33 (1987) 323-328.
D. A. Gough, J. Y. Luciasana and P. H. S. Tse, Two-dimensional enzyme electrode sensor for glucose, Anal. Chem., 57 (1985) 2351-2357. J. L. Romette, B. Froment and D. Thomas, Glucase-oxidase electrode. Measurements of glucose in samples exhibiting high variability in oxygen content, Clin. Chirn. Acta, 9.5 (1979) 249-253. D. J. Claremont, I. E. Sambrook, C. Penton and J. C. Pickup, Subcutaneous implantation of a ferrocene-mediated glucose sensor in pigs, Diabetologia, 29 (1986) 817-821.
I3 (1988) 79-86.
12 Y. Ikariyama, S. Yamauchi, T. Yukiashi and H. Ushioda, Micro-enzyme electrode prepared on platinized platinum, Anal. Lett., 20 (9) (1987) 14071416.
13 Y. Ikariyama, S. Yamauchi, T. Yukiashi and H. Ushioda, Electrochemical fabrication of amperometric microenzyme sensor, J. Electrochem. Sot. I36 (3) ( 1989) 702-706.
14 R. T. Morrison and R. N. Boyd, Organic Chemistry, Allyn and Bacon, Inc., Boston, MA, 3rd edn., 1973, p. 1137. I5 J. J. Mastrototaro, K. W. Johnson, R. J. Morff, D. Lipson, C. C. Andrew and D. J. Allen, An electroenzymatic glucose sensor fabricated on a flexible substrate, Sensors and Actuators B, 5 (1991) 139-144.
Biography Kirk Johnson obtained B.S. and M.A. degrees in chemistry from the University of South Dakota, Vermillion in 1984 and 1986, respectively. He is a chemist in the Medical Device and Diagnostics Division of Eli Lilly and Company. His scientific interests include biosensors, electroanalytical chemistry and electrogenerated chemiluminescence reactions.