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Chip biosensors on thin-film metal electrodes
Sensors and Actuators
B, 4 (1991) 287-291
287
Chip Biosensors on Thin-film Metal Electrodes R. HINTSCHE,
B. MijLLER,
Central Institute of Molecular
I. DRANSFELD, Biology,
U. WOLLENBERGER
Robert-Riissle-Strasse
and F. SCHELLER
IO, I I I5 Berlin (Germany)
B. HOFFMANN University
of Chemnitz,
Chemnitz
(Germany)
Abstract A set of amperometric, impedimetric, potentiometric and reference thin-film metal electrodes (TFMEs) in silicon technology, which may be combined in different shapes and sizes, has been constructed as transducers for biosensors. The chips inserted in test strips or flow-through devices are covered by the enzyme included in surface-crosslinked polyurethane layers. Improved selectivity for glucose, lactose, lysine and urea sensors as well as reduced protein adsorption from real blood samples have been achieved by using chemically modified polymer surfaces. Introduction Because of the limited lifetime of the biocomponent, transducers for biosensors should be cheap, easily exchangeable,. and should permit simple enzyme-immobilization procedures to be used. Several thin- and thick-film metal electrodes, sometimes with integrated reference electrodes, are proposed [l-4]. Most of the described enzyme immobilization procedures at planar electrodes or field-effect transistors use multistep and/or aqueous techniques [5--S]. Apart from choosing different supporting polymers, there are only a few possibilities to adjust the properties of the coatings. The objective of this paper is to demonstrate that a low-cost construction set of simple electrodes may simplify the handling and application of thin-film metal electrodes 0925-4005/91/$3.50
(TFMEs) for biosensors. The described enzyme immobilization allows large-scale manufacturing to be achieved and the parameters of the sensors to be adjusted by a crosslinking procedure. Thin-film Metal Electrodes Different metal layer combinations for forming TFMEs for the potentiometric, amperometric or impedimetric detection of products of the analyte conversion within enzyme-containing layers are shown in Fig. 1. The metal films are deposited by vapour or an electron beam onto common oxidized silicon wafers using metal masks. This procedure allows an inexpensive combination of the different types of electrodes as well as the size and number of electrodes on one chip. The silver chloride layer of the reference electrode was also deposited in the dry process. The antimony oxide is formed from antimony by use of the complete biosensor in water. The mechanical stability of the noble metal layers allows surface mounting technologies to be used for the chips without wire bonding and, furthermore, repeated covering and removing of the enzyme layers. The TFMEs omperometrlc ~mpedimetnc
PHsensitive
Fig. 1. Cross sections of several thin-film metal electrodes. 0
Elsevier
Sequoia/Printed
in The Netherlands
288
Fig. 2. The TFME device: left, test strip; right, flow-through cell. (1, mask; 2, SiO,; 3, Si wafer: 4. working 5, counter electrode; 6, reference electrode; 7, connector; 8, polymer cap.)
are also stable in water for more than one year and in electrolytes at potentials between + 1.5 V and - 1.5 V. The potential of the antimony electrodes [6] shows a stable pH dependence of 56 mV/ pH between pH 2 and pH 9 over a period of 30 days. As shown in Fig. 2 (left), the chips are surface mounted on printed circuit boards isolated by epoxy resin and 0.2 mm thick masks of glass epoxy laminate. In the flowthrough version a gap containing tubing connectors is mounted by epoxy resin on the device (Fig. 2 (right)). The cap and 0.3 mm thick rubber membrane packing with a hole on the active area may also be pressed onto the TFME device. This allows convenient handling and quick exchange of enzyme-covered chips. The active area of the working electrodes has been varied between 0.7 and 7 mm*. The reference TFMEs of 7 mm2 were on the same or on a separate chip using buffers that contained 0.1 mol/l chloride ions. If 100 mm2 silver chloride-covered silver plates or saturated calomel electrodes are combined with the different working TFMEs instead of the reference TFMEs, no differences in the biosensor responses were observed. Ampero-
L
electrodes,
uR
Fig. 3. Scheme of interdigital Pt-TFME for impedance measurements. (Chip size, 4 x 4 mm; interdigital distance, 50 pm.)
metric measurements were carried out in two- or three-electrode mode using lo30 mm2 Pt-TFME as counter electrodes. For impedimetric measurements, interdigital PtTFMEs have been produced by electron beam lithography (Fig. 3). These measurements were carried out in difference mode (500 Hz a.c.) using both an active and an inactive enzyme-covered pair of electrodes.
Polyurethane Enzyme Membrane
The covering enzyme membranes of TFMEs must adhere well at the sensor surface and exhibit a minimal diffusion resistance for substrates and products but a
289
maximal one for interferences. The final biosensors had to be stable during storage and in long-term use. Furthermore, the technologies used should be compatible with mass production. Our basic procedure for the coverage of a TFME by a crosslinked polyurethane enzyme layer, which has also been applied to cover field-effect transistors, has been described earlier [8- 111. The enzyme molecules are included in an elastic and strongly adhesive polyurethane (PU) film. The enzyme dispersion is applied onto the active area of the TFMEs in a diluted organic PU solution by simple dropping, dipping or spinning procedures. The highly porous PU membrane formed is unable to retain the enzyme molecules in water and have therefore been crosslinked. In the following general formula reaction with polyisothe crosslinking cyanates, which reduces the pore sizes of the PU membrane and its swelling in water [7], is shown (n = 3; R, aromatic residue; Ri, enzyme or polyurethane chain; X, amino or hydroxyl group).
1.0 PU
PU
sulfumlenafion
1
PUlTDl
5-
10
19
Fig. 4. Influence of surface polyurethane groups at PU/GOD membranes on the sensitivity to glucose, ascorbic acid and paracetamol compared with PU/ GOD membranes covered with polyestersulfonate or nafion (1 15 = relative number of groups by crosslinking.)
NHCONHR,
/ R-( NCO), + 2RiX + HZO--*R-NHCOORi \ NHCOOH
When glucose oxidase (GOD) is used as the biocatalyst, the resulting enzyme membrane can be stored dry for more than three months and used for more than 300 days for glucose monitoring. The sensor stored in buffer at 280 K may be dried more than ten times between glucose measurements without loss of function. Diphenylmethane diisocyanate or tolylene diisocyanate as crosslinking agents produce different numbers of additional urethane groups at the PU membrane depending on their concentration and the reaction time. The amount of partially charged urethane groups markedly influences the diffusion resistance for analytes and interfering substances (Fig. 4). The diffusional properties and the mechanical stability of the organic layer are better than those of nafion [ 121 or polyestersulfonate [ 131 covering.
TFME glucose sensors which are crosslinked by diphenylmethane diisocyanate exhibited a continuous loss of sensitivity in flow injection analysis (FIA) measurements of serum and blood samples. When chemically modified crosslinking reagents, e.g., adducts of tolylene diisocyanate and low molecular weight polyalcohols, are used as crosslinking reagents, stable responses and standard deviations below 2% are obtained. This suggests that the electrode fouling of this PU enzyme membrane is prevented.
Parameters of the Developed Biosensors
Different biosensors for glucose, lactate, lysine and urea have been constructed using the PU crosslinking method (Table 1). It is obvious that the application of the PU crosslinking method to potentiometric TFMEs (Fig. 5) extends the measuring range, e.g., for
290
TABLE TFME
1. Functional type
Amperom. Pt (test strip) Amperom. Pt (flow through) Impedim. Pt (test strip) Potentiom. Sb (test strip) Potentiom. Sb (test strip) Amperom. Pt (test strip) Amperom. Pt (test strip)
parameters
of different
Analyte
types of biosensors
based on TFMEs
Range [ mmol/l]
Sensitivity
glucose
0.05-6
glucose (in sera) glucose
0.1-50 (FIA) 0.5-2
70 mA/mol (0.9 mm2) 50 mA/mol (0.7 mm’) 70 mF/mol
glucose
0.5-20
41 mV/lg c
30
urea
0.2-5
- 44 mV/lg c
10
lactate
0.0050-0.1
lysine
0.005-0.040
1w l&l (mv)
r
Fig. 5. Calibration curves of potententiometric glucose and urea TFME chips (glucose oxidase and urease in crosslinked PU at pH-sensitive Sb-TFME.)
glucose, as compared with the amperometric device. The limited lifetime of about 30 days for the antimony electrodes should be tolerable with respect to the low-cost sensor. The lifetime of the amperometric glucose test strips of more than 300 days is longer than necessary for their practical use. A lifetime of six weeks for the glucose flow-through chips in flow injection analysis allows more than the practical limit of 1000 measurements to be achieved. During six weeks of measuring and storing at room temperature, no fouling or loss of function has been ‘observed. The time for 95% response of the amperometric PU/TFME devices, reported to be about 2 s [8], has been lowered to less than
Lifetime
[dl
8 mA/mol (7 mm2) 24 mA/mmol (7 mm2)
> 300 > 50 > 50
>
10
>
3
1 s by optimizing the covering procedure. For the potentiometric and also the impedimetric TFMEs, the response time increases to about 10 s using identical PU enzyme membranes. The fast responses of the amperometric biosensors indicate that the crosslinking occurs only near to the surface of the PU membrane and the residual membrane retains its high porosity.
References . A. Turner, I. Karube and G. S. Wilson, in I. Karube (ed.), Biosensors, Vol. 2, Oxford University Press, 1987, pp. 471-480. T. Murakami, S. Nakamoto, J. Kimura, T. Kuriyame and I. Karube, A micro planar amperometric glucose sensor using an ISFET as a reference electrode, Anal. Lett, 19 ( 1986) 1973- 1986. S. Gernet, M. Koudelka and N. F. de Rooij, A planar glucose anzyme electrode, Sensors and Actuators, I7 (1989) 537-540. S. J. Updike, M. C. Shult, C. C. Capelli, D. V. Heimburg, R. K. Rhoses, N. Joseph-Tipton, B. Anderson and D. D. Koch, Laboratory evaluation of new reusable blood glucose sensor, Diabetes Care, 10 (1988) 801-806. S. Shiono, Y. Hanazato and M. Nakako, Urea and glucose sensors based on ion sensitive field effect transistor with photolithographically patterned enzyme membrane, Anal. Sci. 2 ( 1986) 517521.
291
6 J. Kimure, N. Ito, T. Kuryama, M. Kikuchi, T. Arai, N. Negishi and Y. Tomita, A novel blood glucose monitoring method an ISFET microbiosensor applied to transcutaneous effusion fluid, J. Electrochem. Sot., 6 ( 1989) 1744 1747. 7 M. Gotoh, E. Tamiya and I. Karube, Polyvinylbutyral resin membrane for enzyme immobilization to an ISFET microbiosensor, J. Mol. Cutal., 37 ( 1986) 133-139. 8 G. Kampfrath and R. Hintsche, Enzyme immobilization at plasmapolymers, Anal. Lett., 22 ( 1989) 2423-2431. 9 R. Hintsche,
G. Neumann, I. Dransfeld, G. Kampfrath, B. Hoffmann and F. Scheller, Polyurethane enzyme membrane for chip biosensors, Anal. Lett. 22 (1989) 21752190.
10 R. Hintsche, I. Dransfeld, F. Scheller, M. T. Pham, W. Hoffmann, J. Heuller and W. Moritz, Integrated difference enzyme sensors using hydrogen and fluoride ion sensitive multigate FETs, Biosensors BioeIectron., 5 (1990) 327-334. II I. Dransfeld, R. Hintsche, W. Moritz, M. T. Pham,
W. Hoffmann and J. Heuller, Bioscnsors for glucose and lactate using fluoride ion sensitive field effect transistors, Anal. Lett., 23 ( 1990) 437-450. 12 D. J. Harrison, R. F. B. Turner and H. P. Baltes, Characterization of perfluorosulfonic acid polymer coated enzyme electrodes and a miniaturized integrated potentiostat for glucose analysis in whole blood, Anal. Chem., 19 (1988) 2004-2007. 13 L. Gorton, personal communication.
B, 4 (1991) 287-291
287
Chip Biosensors on Thin-film Metal Electrodes R. HINTSCHE,
B. MijLLER,
Central Institute of Molecular
I. DRANSFELD, Biology,
U. WOLLENBERGER
Robert-Riissle-Strasse
and F. SCHELLER
IO, I I I5 Berlin (Germany)
B. HOFFMANN University
of Chemnitz,
Chemnitz
(Germany)
Abstract A set of amperometric, impedimetric, potentiometric and reference thin-film metal electrodes (TFMEs) in silicon technology, which may be combined in different shapes and sizes, has been constructed as transducers for biosensors. The chips inserted in test strips or flow-through devices are covered by the enzyme included in surface-crosslinked polyurethane layers. Improved selectivity for glucose, lactose, lysine and urea sensors as well as reduced protein adsorption from real blood samples have been achieved by using chemically modified polymer surfaces. Introduction Because of the limited lifetime of the biocomponent, transducers for biosensors should be cheap, easily exchangeable,. and should permit simple enzyme-immobilization procedures to be used. Several thin- and thick-film metal electrodes, sometimes with integrated reference electrodes, are proposed [l-4]. Most of the described enzyme immobilization procedures at planar electrodes or field-effect transistors use multistep and/or aqueous techniques [5--S]. Apart from choosing different supporting polymers, there are only a few possibilities to adjust the properties of the coatings. The objective of this paper is to demonstrate that a low-cost construction set of simple electrodes may simplify the handling and application of thin-film metal electrodes 0925-4005/91/$3.50
(TFMEs) for biosensors. The described enzyme immobilization allows large-scale manufacturing to be achieved and the parameters of the sensors to be adjusted by a crosslinking procedure. Thin-film Metal Electrodes Different metal layer combinations for forming TFMEs for the potentiometric, amperometric or impedimetric detection of products of the analyte conversion within enzyme-containing layers are shown in Fig. 1. The metal films are deposited by vapour or an electron beam onto common oxidized silicon wafers using metal masks. This procedure allows an inexpensive combination of the different types of electrodes as well as the size and number of electrodes on one chip. The silver chloride layer of the reference electrode was also deposited in the dry process. The antimony oxide is formed from antimony by use of the complete biosensor in water. The mechanical stability of the noble metal layers allows surface mounting technologies to be used for the chips without wire bonding and, furthermore, repeated covering and removing of the enzyme layers. The TFMEs omperometrlc ~mpedimetnc
PHsensitive
Fig. 1. Cross sections of several thin-film metal electrodes. 0
Elsevier
Sequoia/Printed
in The Netherlands
288
Fig. 2. The TFME device: left, test strip; right, flow-through cell. (1, mask; 2, SiO,; 3, Si wafer: 4. working 5, counter electrode; 6, reference electrode; 7, connector; 8, polymer cap.)
are also stable in water for more than one year and in electrolytes at potentials between + 1.5 V and - 1.5 V. The potential of the antimony electrodes [6] shows a stable pH dependence of 56 mV/ pH between pH 2 and pH 9 over a period of 30 days. As shown in Fig. 2 (left), the chips are surface mounted on printed circuit boards isolated by epoxy resin and 0.2 mm thick masks of glass epoxy laminate. In the flowthrough version a gap containing tubing connectors is mounted by epoxy resin on the device (Fig. 2 (right)). The cap and 0.3 mm thick rubber membrane packing with a hole on the active area may also be pressed onto the TFME device. This allows convenient handling and quick exchange of enzyme-covered chips. The active area of the working electrodes has been varied between 0.7 and 7 mm*. The reference TFMEs of 7 mm2 were on the same or on a separate chip using buffers that contained 0.1 mol/l chloride ions. If 100 mm2 silver chloride-covered silver plates or saturated calomel electrodes are combined with the different working TFMEs instead of the reference TFMEs, no differences in the biosensor responses were observed. Ampero-
L
electrodes,
uR
Fig. 3. Scheme of interdigital Pt-TFME for impedance measurements. (Chip size, 4 x 4 mm; interdigital distance, 50 pm.)
metric measurements were carried out in two- or three-electrode mode using lo30 mm2 Pt-TFME as counter electrodes. For impedimetric measurements, interdigital PtTFMEs have been produced by electron beam lithography (Fig. 3). These measurements were carried out in difference mode (500 Hz a.c.) using both an active and an inactive enzyme-covered pair of electrodes.
Polyurethane Enzyme Membrane
The covering enzyme membranes of TFMEs must adhere well at the sensor surface and exhibit a minimal diffusion resistance for substrates and products but a
289
maximal one for interferences. The final biosensors had to be stable during storage and in long-term use. Furthermore, the technologies used should be compatible with mass production. Our basic procedure for the coverage of a TFME by a crosslinked polyurethane enzyme layer, which has also been applied to cover field-effect transistors, has been described earlier [8- 111. The enzyme molecules are included in an elastic and strongly adhesive polyurethane (PU) film. The enzyme dispersion is applied onto the active area of the TFMEs in a diluted organic PU solution by simple dropping, dipping or spinning procedures. The highly porous PU membrane formed is unable to retain the enzyme molecules in water and have therefore been crosslinked. In the following general formula reaction with polyisothe crosslinking cyanates, which reduces the pore sizes of the PU membrane and its swelling in water [7], is shown (n = 3; R, aromatic residue; Ri, enzyme or polyurethane chain; X, amino or hydroxyl group).
1.0 PU
PU
sulfumlenafion
1
PUlTDl
5-
10
19
Fig. 4. Influence of surface polyurethane groups at PU/GOD membranes on the sensitivity to glucose, ascorbic acid and paracetamol compared with PU/ GOD membranes covered with polyestersulfonate or nafion (1 15 = relative number of groups by crosslinking.)
NHCONHR,
/ R-( NCO), + 2RiX + HZO--*R-NHCOORi \ NHCOOH
When glucose oxidase (GOD) is used as the biocatalyst, the resulting enzyme membrane can be stored dry for more than three months and used for more than 300 days for glucose monitoring. The sensor stored in buffer at 280 K may be dried more than ten times between glucose measurements without loss of function. Diphenylmethane diisocyanate or tolylene diisocyanate as crosslinking agents produce different numbers of additional urethane groups at the PU membrane depending on their concentration and the reaction time. The amount of partially charged urethane groups markedly influences the diffusion resistance for analytes and interfering substances (Fig. 4). The diffusional properties and the mechanical stability of the organic layer are better than those of nafion [ 121 or polyestersulfonate [ 131 covering.
TFME glucose sensors which are crosslinked by diphenylmethane diisocyanate exhibited a continuous loss of sensitivity in flow injection analysis (FIA) measurements of serum and blood samples. When chemically modified crosslinking reagents, e.g., adducts of tolylene diisocyanate and low molecular weight polyalcohols, are used as crosslinking reagents, stable responses and standard deviations below 2% are obtained. This suggests that the electrode fouling of this PU enzyme membrane is prevented.
Parameters of the Developed Biosensors
Different biosensors for glucose, lactate, lysine and urea have been constructed using the PU crosslinking method (Table 1). It is obvious that the application of the PU crosslinking method to potentiometric TFMEs (Fig. 5) extends the measuring range, e.g., for
290
TABLE TFME
1. Functional type
Amperom. Pt (test strip) Amperom. Pt (flow through) Impedim. Pt (test strip) Potentiom. Sb (test strip) Potentiom. Sb (test strip) Amperom. Pt (test strip) Amperom. Pt (test strip)
parameters
of different
Analyte
types of biosensors
based on TFMEs
Range [ mmol/l]
Sensitivity
glucose
0.05-6
glucose (in sera) glucose
0.1-50 (FIA) 0.5-2
70 mA/mol (0.9 mm2) 50 mA/mol (0.7 mm’) 70 mF/mol
glucose
0.5-20
41 mV/lg c
30
urea
0.2-5
- 44 mV/lg c
10
lactate
0.0050-0.1
lysine
0.005-0.040
1w l&l (mv)
r
Fig. 5. Calibration curves of potententiometric glucose and urea TFME chips (glucose oxidase and urease in crosslinked PU at pH-sensitive Sb-TFME.)
glucose, as compared with the amperometric device. The limited lifetime of about 30 days for the antimony electrodes should be tolerable with respect to the low-cost sensor. The lifetime of the amperometric glucose test strips of more than 300 days is longer than necessary for their practical use. A lifetime of six weeks for the glucose flow-through chips in flow injection analysis allows more than the practical limit of 1000 measurements to be achieved. During six weeks of measuring and storing at room temperature, no fouling or loss of function has been ‘observed. The time for 95% response of the amperometric PU/TFME devices, reported to be about 2 s [8], has been lowered to less than
Lifetime
[dl
8 mA/mol (7 mm2) 24 mA/mmol (7 mm2)
> 300 > 50 > 50
>
10
>
3
1 s by optimizing the covering procedure. For the potentiometric and also the impedimetric TFMEs, the response time increases to about 10 s using identical PU enzyme membranes. The fast responses of the amperometric biosensors indicate that the crosslinking occurs only near to the surface of the PU membrane and the residual membrane retains its high porosity.
References . A. Turner, I. Karube and G. S. Wilson, in I. Karube (ed.), Biosensors, Vol. 2, Oxford University Press, 1987, pp. 471-480. T. Murakami, S. Nakamoto, J. Kimura, T. Kuriyame and I. Karube, A micro planar amperometric glucose sensor using an ISFET as a reference electrode, Anal. Lett, 19 ( 1986) 1973- 1986. S. Gernet, M. Koudelka and N. F. de Rooij, A planar glucose anzyme electrode, Sensors and Actuators, I7 (1989) 537-540. S. J. Updike, M. C. Shult, C. C. Capelli, D. V. Heimburg, R. K. Rhoses, N. Joseph-Tipton, B. Anderson and D. D. Koch, Laboratory evaluation of new reusable blood glucose sensor, Diabetes Care, 10 (1988) 801-806. S. Shiono, Y. Hanazato and M. Nakako, Urea and glucose sensors based on ion sensitive field effect transistor with photolithographically patterned enzyme membrane, Anal. Sci. 2 ( 1986) 517521.
291
6 J. Kimure, N. Ito, T. Kuryama, M. Kikuchi, T. Arai, N. Negishi and Y. Tomita, A novel blood glucose monitoring method an ISFET microbiosensor applied to transcutaneous effusion fluid, J. Electrochem. Sot., 6 ( 1989) 1744 1747. 7 M. Gotoh, E. Tamiya and I. Karube, Polyvinylbutyral resin membrane for enzyme immobilization to an ISFET microbiosensor, J. Mol. Cutal., 37 ( 1986) 133-139. 8 G. Kampfrath and R. Hintsche, Enzyme immobilization at plasmapolymers, Anal. Lett., 22 ( 1989) 2423-2431. 9 R. Hintsche,
G. Neumann, I. Dransfeld, G. Kampfrath, B. Hoffmann and F. Scheller, Polyurethane enzyme membrane for chip biosensors, Anal. Lett. 22 (1989) 21752190.
10 R. Hintsche, I. Dransfeld, F. Scheller, M. T. Pham, W. Hoffmann, J. Heuller and W. Moritz, Integrated difference enzyme sensors using hydrogen and fluoride ion sensitive multigate FETs, Biosensors BioeIectron., 5 (1990) 327-334. II I. Dransfeld, R. Hintsche, W. Moritz, M. T. Pham,
W. Hoffmann and J. Heuller, Bioscnsors for glucose and lactate using fluoride ion sensitive field effect transistors, Anal. Lett., 23 ( 1990) 437-450. 12 D. J. Harrison, R. F. B. Turner and H. P. Baltes, Characterization of perfluorosulfonic acid polymer coated enzyme electrodes and a miniaturized integrated potentiostat for glucose analysis in whole blood, Anal. Chem., 19 (1988) 2004-2007. 13 L. Gorton, personal communication.