CA2163816C - Electrochemical enzyme biosensor - Google Patents
Electrochemical enzyme biosensor Download PDFInfo
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- CA2163816C CA2163816C CA002163816A CA2163816A CA2163816C CA 2163816 C CA2163816 C CA 2163816C CA 002163816 A CA002163816 A CA 002163816A CA 2163816 A CA2163816 A CA 2163816A CA 2163816 C CA2163816 C CA 2163816C
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- electrochemical
- enzyme biosensor
- biosensor according
- electrochemical enzyme
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- C—CHEMISTRY; METALLURGY
- C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
- C12Q—MEASURING OR TESTING PROCESSES INVOLVING ENZYMES, NUCLEIC ACIDS OR MICROORGANISMS; COMPOSITIONS OR TEST PAPERS THEREFOR; PROCESSES OF PREPARING SUCH COMPOSITIONS; CONDITION-RESPONSIVE CONTROL IN MICROBIOLOGICAL OR ENZYMOLOGICAL PROCESSES
- C12Q1/00—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions
- C12Q1/001—Enzyme electrodes
- C12Q1/004—Enzyme electrodes mediator-assisted
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- Y—GENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
- Y10—TECHNICAL SUBJECTS COVERED BY FORMER USPC
- Y10S—TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
- Y10S435/00—Chemistry: molecular biology and microbiology
- Y10S435/817—Enzyme or microbe electrode
Abstract
In the case of the electrochemical enzyme biosensor comprising noble metal electrodes, an electrochemical oxidation of pyridine nucleotides takes place, in particular of NADH. The noble metal electrodes have a microtexture rich in surface pores and having catalytic properties, as a result of which the overpotential required for electrochemical oxidation of the pyridine nucleotides is reduced. Owing to these microtextures, the surface area of an electrode is from 3 to 10 times larger than its geometric surface area.
Description
21638 ~6 ELECTROCHEMICAL ENZYME BIOSENSOR
The invention is based on an electrochemical enzyme biosensor which comprises noble metal electrodes and with 5 which an electrochemical oxidation of pyridine nucleo-tides, in particular of NADH takes place.
Numerous enzyme analysis methods in clinical diag-nostics and in environmental, process and food analysis 10 are based on reactions which are catalyzed by enzymes of the dehydrogenase class. This involves utilizing the di-rect proportionality between the concentration of the substrate to be analyzed on the one hand and the concen-tration of the coenzyme NADH formed or consumed on the 15 other hand. Most of these methods take advantage of op-tical properties of the molecule NADH, which differs, for example, from its corresponding oxidized form NAD+ by a characteristic absorption of light having a wavelength of 360 nm. In the case of coloured or turbid analytes, 20 these spectrophotometric methods will largely fail or at least require a laborious and time-consuming pretreatment of the samples. This can be circumvented by electro-chemical detection of NADH. Moreover, direct, stable coupling of electrodes'and enzyme to form electrochemical 25 biosensors is possible. The drawback of the electro-chemical method has hitherto been the high overpotential required for oxidizing NADH. This also results in the detection of a number of further substances which are likewise present in the analyte, so that unduly high 30 analysis values are normally obtained, which subsequently have to be corrected. Moreover the electrochemical oxi-dation of NADH at high overpotentials leads, via radical intermediates, to dimers and oligomers of NADH which per-MSE #2105 manently c:oat the electrode surface and cause lasting damage (so-called electrode fouling, cf. W. J. Blaedel &
R. A. Jenk:ins, Study of the electrochemical oxidation of reduced nicotinamide adenine dinucleotide, Anal. Chem. 47 5 (1975) 13?~7-1343). To circumvent this phenomenon, the analyte is often admixed with catalytic amounts of low-molecular weight molecules which are able to undergo re-dox reactions and which effect the transport of the elec-trons from NADH to the electrode ( cf . , amongst many oth-10 ers, L. Gorton, Chemically modified electrodes for the electrocatalytic oxidation of nicotinamide coenzymes, J.
Chem. Soc. Faraday Trans. 82 (1986) 1245-1258). These so-called mediators are in part immobilized on the elec-trode, although in view of the restriction of the media-15 for having to remain mobile in order to be effective, stable binding of the mediator to the electrode is often not possible. The mediator diffuses into the analyte, to the considerable detriment of the stability of the sen-sors.
A method described in U.S. Patent Specification 5,240,571 employs a coupling reagent capable of undergo-ing redox reactions, which is added to the analyte and, with NADH, forms an electroactive compound which is oxi-25 dized at 7Low overpotentials. U.S. Patent Specification 5,122,457 discloses that NADH can be oxidized quantita-tively on porous carbon electrodes and graphite elec-trodes coated with platinum or palladium, the potential employed not being specified.
The object of the invention is to improve an elec-trochemical enzyme biosensor comprising noble metal elec-MSE #2105 ________ 2163816 trodes in such a way, without adding a mediator to the analyte and without any chemical modification of the electrode, that any interfering substances which may be present in the analyte and which are likewise electro-chemically oxidizable are largely unable to enter into any electrode reaction. In addition to the higher selec-tivity thus obtained, the stability of the transductor electrodes and thus the operational stability of the sen-sors fabricated therefrom is to be enhanced.
This object is achieved, according to the invention, and starting from an electrochemical enzyme biosensor which comprises noble metal electrodes and with which an electrochennical oxidation of pyridine nucleotides takes place, by the noble metal electrodes having a microtex-ture rich in surface pores and having catalytic proper-ties, as a result of which the overpotential required for electrochennical oxidation of the pyridine nucleotides is reduced. The microtexture rich in surface pores is caused by numerous microscopically small peaks and_cra-ters, so that the surface area of the electrodes is con-siderably increased. The peaks consist of noble metal particles or of particle clusters which project from the electrode aurface and have a mean radius of curvature of from 0.05 um to 4 Nm, preferably from 0.5 pm to 2 Nm.
This microroughness increases the surface area of an electrode by approximately a factor of from 3 to 10, preferably from 4 to 7 (compared with the geometric sur-face area),, which can be demonstrated by electron micros-copy and irnpedance spectroscopy measurements.
MSE #2105 This special microtexture, rich in surface pores, of the noble metal electrodes results in the electrodes hav-ing catalytic properties which manifest themselves in a drastic reduction of the overpotential for the oxidation 5 of pyridine nucleotides, especially of NADH. It was found that the overpotential required for the anodic oxi-dation of NADH to NAD+ is reduced to from 0 to 200 mV, preferably to from 100 to 150 mV, against the SCE and is thus at least 50% below the values normally found.
Particularly good results are achieved if the micro-porous noble metal electrodes are made of gold.
Immobilization of the selectivity-determining en-zymes can preferably be effected by a membrane containing the enzyme being applied to the working electrode. Adhe-sion of the membrane on the planar electrode is promoted by the above-described microroughness of the electrode.
The use, in particular, of membranes made from an aqueous 20 poly(vinylacetate) dispersion, such as are described, e.g., in c3erman Patent Specification 40 27 728, allows stable, selective, highly sensitive biosensors to be ob-tained.
25 According to a further development, the microporous noble metal electrodes are modified by coating them with electropol~ymerizable, conductive polymers, in particular with polypyrrole and poly(methylene blue). This modifi-cation can be utilized for the immobilization of enzymes 30 and other selectivity-conferring biocomponents. In the process, the disperse surface of the noble metal elec-trodes enables homogeneous growth of the polymer during MSE #2105 2'638 16 the electropolymerization, since many reactive sites, uniformly distributed over the entire macroscopic elec-trode surface, are generated.
5 If the electrochemical polymerization is carried out in the presence of an enzyme, this is immobilized by be-ing physically trapped in the polymer during the growth of the latter in front of the electrode. The polymer skeleton i:irstly serves to immobilize the enzyme, but 10 also has the function of a sieve which prevents large molecules, e.g. proteins, from reaching the electrode surface, dating it and thus at least reversibly damaging it. Furthermore, the polymer, if it has redox sites, can act as a catalytic mediator for the electron transfer 15 from the enzyme or the coenzyme to the metal electrode.
In the case of poly(methylene blue), the amplitude of the amperometric signal can therefore be distinctly increased (factor 10), which has a beneficial effect on the signal-to-noise ratio and thus results in a lower detection 20 limit and an enhanced sensitivity for the substances to be determined .
Alternatively, enzymes or other biocomponents are immobilized directly on the electrodes via reactive pen-25 dent groups. This permits the construction of biosensors having extremely short response times, compared with mem-brane biosensors, since the component to be measured does not have to diffuse through a membrane. Some enzymes such as, E~.g., glucose oxidase, are glycoproteins whose 30 polysaccharide envelope is suitable for chemical linking to bifunci~ional spacer molecules. These spacers are bound to t:he sugar residues of the glycoprotein via one MSE #2105 ~16~g 16 of their i:wo functional groups, e.g. by an amino func-tion, whereas the other functional terminal group, e.g. a thiol group, enters into a stable bond with the surface of the gold electrodes. In the case of glucose oxidase, 5 the enzyme lyophilisate is dissolved in a carbonate buffer (pH 8.1) and after the addition of one per cent strength e~thanolic 2,4-dinitrofluorobenzene solution, is admixed with sodium periodate solution (60 mM), whereupon a so-called periodate cleavage of the sugar component of 10 the enzyme takes place. Reactive carbonyl groups gener-ated in th.e process are then, after a dialytic purifica-tion step, coupled with a bifunctional thio, e.g. with cystamine, via its amino groups. The modified enzyme ob-tained is stably bound to the gold surface via the thiol 15 groups of the cystamine radical.
The above-described microtextures rich in surface pores are produced, in the course of the fabrication of gold electrodes, by a chemically inert base being coated, 20 by means ~~f screen printing, with a paste comprising a gold powder and a polymeric binder, and the coated base then beings passed through a furnace having incrementally gradated firing temperature zones, the temperature of the first firing zone being from 300°C to 400°C and that of 25 the last firing zone being from 800°C to 950°C. The residence time of the coated base in this process is preferably approximately 3 min to 8 min.
The invention provides the following advantages:
It was found that the overpotential for the anodic oxidation of NADH to NAD' on the noble metal electrodes MSE #2105 2~s3$ ~s -fabricated by a thick-film technique by means of screen printing is reduced to such an extent that - in contrast to conventional noble metal electrodes, which require an overpotent~:al of +600 mV - for the purpose of amperomet-5 ric measurements of NADH concentrations, overpotentials of less than +200 mV against a KC1-saturated aqueous calomel reference electrode (SCE) will suffice. The re-duction in the overpotential results in any interfering substances cooxidizable at high overpotentials, which may 10 be present in the liquid analyte, no longer undergoing electrochemical reactions and no longer falsifying the result of the analysis. The noble metal electrodes de-scribed are utilized especially as transductors for elec-trochemical enzyme biosensors based on dehydrogenases.
15 The main fields of application of these sensors are clinical diagnostics and environmental, process and food analysis.
The invention is explained below in more detail with 20 reference to a fabrication example and to illustrative measurements.
Fabrication Example 25 A gold powder having a mean particle size of 1.5 Nm and a sizE~ distribution of from 1 pm to 3.5 Nm is inti-mately mixed with a nitrocellulose-containing polymeric binder. The gold paste thus prepared is then applied by screen-printing to a ceramic base plate having the dimen-30 sions 41 a 41 x 0.5 mm3. Then the coated base plate is run, at a speed of 8 cm/min, through a furnace comprising MSE #2105 ~~s3a ~s six firing zones, each 60 cm in length, having the fol-lowing temperature profile:
1st zone 336C
:2nd zone 685C
3rd zone 872C
4th zone 905C
5th zone 857C
10 The particle size distribution of the gold particles in the screen printing paste employed and their radii de-termine th.e subsequent surface properties and the poros-ity of thE~ gold electrodes. Additionally, minute explo-sions which take place during the pyrolysis of the poly-15 mer fraction result in sharp-edged craters forming on the gold surface which likewise affect the surface proper-ties. The minute explosions are caused by polymer layers situated further down being pyrolyzed so rapidly that the gas diffusion of the layers lying above them is not suf-20 ficient to bring down the resulting pressure. By means of comparative impedance spectroscopy measurements of the actual surface area of conventional, polished gold elec-trodes an<i of gold electrodes prepared by a thick-film technique it was shown that the latter, given an identi-25 cal geomei~ric surface area, have six times as large an actual surface area.
Illustrative measurements according to Fig. 1 to 4.
30 In the figures, MSE #2105 ~1 s38 ~ s Figure 1 shows the cyclic voltammograms of (a) a microporous thick-film gold electrode and (b) a commercially available planar gold electrode polished to mirror brightness (EG & G) in buffered solutions with (----) and without ( - ) the addition of NADH.
Figure 2 shows the long-term stability of a micro porous gold electrode during the measure ment of NADH.
Figure 3 shows the effect of interfering substances in the case of conventional potentials and potentials made possible by the invention, for the oxidation of NADH.
Figure 4 shows the concentration-current character-istic of a glucose biosensor which, em-ploying the electrode according to the in-vention as a transductor, operates at low polarization voltages.
Figure 1 illustrates the catalytic activity of the electrode:c according to the invention with respect to the anodic oxidation of NADH. In the course of the cyclic voltammogram there appears, in the case of the thick-film gold electrode (a) in the presence of NADH (1), starting at approximately +100 mV (against SCE), a distinct anodic current which is caused by the oxidation of NADH. With the NADH-free buffer (2) this current is absent. With conventional electrodes (b), such behaviour cannot be ob-served under identical experimental conditions (3, 4).
The cyclic voltammograms were recorded in 0.1 M sodium MSE #2105 ~! ~ g 38 1 6 _ ___ phosphate buffer of pH 7.0, which contained 0.1 M sodium perchlorate as the supporting electrolyte. The voltage rate was 10 mV/sec. The NADH concentration was 0 mM (2, 4) and 5 niM (1, 3), respectively.
The thick-film gold electrodes according to the in-vention wE~re incorporated into a flow injection system for detecting 0.5 mM NADH and were operated at +145 mV
(against SCE). Figure 2 shows the change of the anodic 10 current against time for the oxidation of NADH. The ar-rows indicate that in each case freshly prepared NADH so-lutions were used. The drop in the current after the in-troduction of a fresh solution therefore mainly results from the thermal decomposition of NADH in solution (0.1 M
sodium phosphate buffer, pH 7.0 with 0.1 M sodium per-chlorate, room temperature ) and is not due, or only to a limited extent, to electrode fouling.
To clarify the advantages of sensor operation at low polarization voltages, potentially interfering substances (1 and 2 mM hydrogen peroxide, 1 and 2 mM acetaminophen (= paracetamol), serum diluted 1:1 and undiluted) were studied together with NADH (1 and 2 mM) in a flow injec-tion system (Fig. 3). In so doing, in one case, the voltage most common hitherto of (a) +555 mV was applied to the electrodes and (b) in another case at the lower potential of +145 mV (both against SCE), which is suffi-cient to oxidize NADH, using tie electrodes according to the invention. Apart from NADH, nothing was indicated at this low ~aotential except, to a small extent, serum com-ponents.
MSE #2105 X1638 1 6 ~ ._ The glucose biosensor whose calibration curve is re-produced in Figure 4 represents a possible application for the electrodes according to the invention. The en-zyme glucose dehydrogenase was immobilized onto the gold 5 surface by matrix inclusion into a polyvinyl acetate) membrane. The calibration curve was recorded in a flow injection system which was operated with 0.1 M sodium phosphate buffer of pH 7.0 and whose injection volume was 100 N1 at a flow rate of 0.6 ml/min.
MSE #2105
The invention is based on an electrochemical enzyme biosensor which comprises noble metal electrodes and with 5 which an electrochemical oxidation of pyridine nucleo-tides, in particular of NADH takes place.
Numerous enzyme analysis methods in clinical diag-nostics and in environmental, process and food analysis 10 are based on reactions which are catalyzed by enzymes of the dehydrogenase class. This involves utilizing the di-rect proportionality between the concentration of the substrate to be analyzed on the one hand and the concen-tration of the coenzyme NADH formed or consumed on the 15 other hand. Most of these methods take advantage of op-tical properties of the molecule NADH, which differs, for example, from its corresponding oxidized form NAD+ by a characteristic absorption of light having a wavelength of 360 nm. In the case of coloured or turbid analytes, 20 these spectrophotometric methods will largely fail or at least require a laborious and time-consuming pretreatment of the samples. This can be circumvented by electro-chemical detection of NADH. Moreover, direct, stable coupling of electrodes'and enzyme to form electrochemical 25 biosensors is possible. The drawback of the electro-chemical method has hitherto been the high overpotential required for oxidizing NADH. This also results in the detection of a number of further substances which are likewise present in the analyte, so that unduly high 30 analysis values are normally obtained, which subsequently have to be corrected. Moreover the electrochemical oxi-dation of NADH at high overpotentials leads, via radical intermediates, to dimers and oligomers of NADH which per-MSE #2105 manently c:oat the electrode surface and cause lasting damage (so-called electrode fouling, cf. W. J. Blaedel &
R. A. Jenk:ins, Study of the electrochemical oxidation of reduced nicotinamide adenine dinucleotide, Anal. Chem. 47 5 (1975) 13?~7-1343). To circumvent this phenomenon, the analyte is often admixed with catalytic amounts of low-molecular weight molecules which are able to undergo re-dox reactions and which effect the transport of the elec-trons from NADH to the electrode ( cf . , amongst many oth-10 ers, L. Gorton, Chemically modified electrodes for the electrocatalytic oxidation of nicotinamide coenzymes, J.
Chem. Soc. Faraday Trans. 82 (1986) 1245-1258). These so-called mediators are in part immobilized on the elec-trode, although in view of the restriction of the media-15 for having to remain mobile in order to be effective, stable binding of the mediator to the electrode is often not possible. The mediator diffuses into the analyte, to the considerable detriment of the stability of the sen-sors.
A method described in U.S. Patent Specification 5,240,571 employs a coupling reagent capable of undergo-ing redox reactions, which is added to the analyte and, with NADH, forms an electroactive compound which is oxi-25 dized at 7Low overpotentials. U.S. Patent Specification 5,122,457 discloses that NADH can be oxidized quantita-tively on porous carbon electrodes and graphite elec-trodes coated with platinum or palladium, the potential employed not being specified.
The object of the invention is to improve an elec-trochemical enzyme biosensor comprising noble metal elec-MSE #2105 ________ 2163816 trodes in such a way, without adding a mediator to the analyte and without any chemical modification of the electrode, that any interfering substances which may be present in the analyte and which are likewise electro-chemically oxidizable are largely unable to enter into any electrode reaction. In addition to the higher selec-tivity thus obtained, the stability of the transductor electrodes and thus the operational stability of the sen-sors fabricated therefrom is to be enhanced.
This object is achieved, according to the invention, and starting from an electrochemical enzyme biosensor which comprises noble metal electrodes and with which an electrochennical oxidation of pyridine nucleotides takes place, by the noble metal electrodes having a microtex-ture rich in surface pores and having catalytic proper-ties, as a result of which the overpotential required for electrochennical oxidation of the pyridine nucleotides is reduced. The microtexture rich in surface pores is caused by numerous microscopically small peaks and_cra-ters, so that the surface area of the electrodes is con-siderably increased. The peaks consist of noble metal particles or of particle clusters which project from the electrode aurface and have a mean radius of curvature of from 0.05 um to 4 Nm, preferably from 0.5 pm to 2 Nm.
This microroughness increases the surface area of an electrode by approximately a factor of from 3 to 10, preferably from 4 to 7 (compared with the geometric sur-face area),, which can be demonstrated by electron micros-copy and irnpedance spectroscopy measurements.
MSE #2105 This special microtexture, rich in surface pores, of the noble metal electrodes results in the electrodes hav-ing catalytic properties which manifest themselves in a drastic reduction of the overpotential for the oxidation 5 of pyridine nucleotides, especially of NADH. It was found that the overpotential required for the anodic oxi-dation of NADH to NAD+ is reduced to from 0 to 200 mV, preferably to from 100 to 150 mV, against the SCE and is thus at least 50% below the values normally found.
Particularly good results are achieved if the micro-porous noble metal electrodes are made of gold.
Immobilization of the selectivity-determining en-zymes can preferably be effected by a membrane containing the enzyme being applied to the working electrode. Adhe-sion of the membrane on the planar electrode is promoted by the above-described microroughness of the electrode.
The use, in particular, of membranes made from an aqueous 20 poly(vinylacetate) dispersion, such as are described, e.g., in c3erman Patent Specification 40 27 728, allows stable, selective, highly sensitive biosensors to be ob-tained.
25 According to a further development, the microporous noble metal electrodes are modified by coating them with electropol~ymerizable, conductive polymers, in particular with polypyrrole and poly(methylene blue). This modifi-cation can be utilized for the immobilization of enzymes 30 and other selectivity-conferring biocomponents. In the process, the disperse surface of the noble metal elec-trodes enables homogeneous growth of the polymer during MSE #2105 2'638 16 the electropolymerization, since many reactive sites, uniformly distributed over the entire macroscopic elec-trode surface, are generated.
5 If the electrochemical polymerization is carried out in the presence of an enzyme, this is immobilized by be-ing physically trapped in the polymer during the growth of the latter in front of the electrode. The polymer skeleton i:irstly serves to immobilize the enzyme, but 10 also has the function of a sieve which prevents large molecules, e.g. proteins, from reaching the electrode surface, dating it and thus at least reversibly damaging it. Furthermore, the polymer, if it has redox sites, can act as a catalytic mediator for the electron transfer 15 from the enzyme or the coenzyme to the metal electrode.
In the case of poly(methylene blue), the amplitude of the amperometric signal can therefore be distinctly increased (factor 10), which has a beneficial effect on the signal-to-noise ratio and thus results in a lower detection 20 limit and an enhanced sensitivity for the substances to be determined .
Alternatively, enzymes or other biocomponents are immobilized directly on the electrodes via reactive pen-25 dent groups. This permits the construction of biosensors having extremely short response times, compared with mem-brane biosensors, since the component to be measured does not have to diffuse through a membrane. Some enzymes such as, E~.g., glucose oxidase, are glycoproteins whose 30 polysaccharide envelope is suitable for chemical linking to bifunci~ional spacer molecules. These spacers are bound to t:he sugar residues of the glycoprotein via one MSE #2105 ~16~g 16 of their i:wo functional groups, e.g. by an amino func-tion, whereas the other functional terminal group, e.g. a thiol group, enters into a stable bond with the surface of the gold electrodes. In the case of glucose oxidase, 5 the enzyme lyophilisate is dissolved in a carbonate buffer (pH 8.1) and after the addition of one per cent strength e~thanolic 2,4-dinitrofluorobenzene solution, is admixed with sodium periodate solution (60 mM), whereupon a so-called periodate cleavage of the sugar component of 10 the enzyme takes place. Reactive carbonyl groups gener-ated in th.e process are then, after a dialytic purifica-tion step, coupled with a bifunctional thio, e.g. with cystamine, via its amino groups. The modified enzyme ob-tained is stably bound to the gold surface via the thiol 15 groups of the cystamine radical.
The above-described microtextures rich in surface pores are produced, in the course of the fabrication of gold electrodes, by a chemically inert base being coated, 20 by means ~~f screen printing, with a paste comprising a gold powder and a polymeric binder, and the coated base then beings passed through a furnace having incrementally gradated firing temperature zones, the temperature of the first firing zone being from 300°C to 400°C and that of 25 the last firing zone being from 800°C to 950°C. The residence time of the coated base in this process is preferably approximately 3 min to 8 min.
The invention provides the following advantages:
It was found that the overpotential for the anodic oxidation of NADH to NAD' on the noble metal electrodes MSE #2105 2~s3$ ~s -fabricated by a thick-film technique by means of screen printing is reduced to such an extent that - in contrast to conventional noble metal electrodes, which require an overpotent~:al of +600 mV - for the purpose of amperomet-5 ric measurements of NADH concentrations, overpotentials of less than +200 mV against a KC1-saturated aqueous calomel reference electrode (SCE) will suffice. The re-duction in the overpotential results in any interfering substances cooxidizable at high overpotentials, which may 10 be present in the liquid analyte, no longer undergoing electrochemical reactions and no longer falsifying the result of the analysis. The noble metal electrodes de-scribed are utilized especially as transductors for elec-trochemical enzyme biosensors based on dehydrogenases.
15 The main fields of application of these sensors are clinical diagnostics and environmental, process and food analysis.
The invention is explained below in more detail with 20 reference to a fabrication example and to illustrative measurements.
Fabrication Example 25 A gold powder having a mean particle size of 1.5 Nm and a sizE~ distribution of from 1 pm to 3.5 Nm is inti-mately mixed with a nitrocellulose-containing polymeric binder. The gold paste thus prepared is then applied by screen-printing to a ceramic base plate having the dimen-30 sions 41 a 41 x 0.5 mm3. Then the coated base plate is run, at a speed of 8 cm/min, through a furnace comprising MSE #2105 ~~s3a ~s six firing zones, each 60 cm in length, having the fol-lowing temperature profile:
1st zone 336C
:2nd zone 685C
3rd zone 872C
4th zone 905C
5th zone 857C
10 The particle size distribution of the gold particles in the screen printing paste employed and their radii de-termine th.e subsequent surface properties and the poros-ity of thE~ gold electrodes. Additionally, minute explo-sions which take place during the pyrolysis of the poly-15 mer fraction result in sharp-edged craters forming on the gold surface which likewise affect the surface proper-ties. The minute explosions are caused by polymer layers situated further down being pyrolyzed so rapidly that the gas diffusion of the layers lying above them is not suf-20 ficient to bring down the resulting pressure. By means of comparative impedance spectroscopy measurements of the actual surface area of conventional, polished gold elec-trodes an<i of gold electrodes prepared by a thick-film technique it was shown that the latter, given an identi-25 cal geomei~ric surface area, have six times as large an actual surface area.
Illustrative measurements according to Fig. 1 to 4.
30 In the figures, MSE #2105 ~1 s38 ~ s Figure 1 shows the cyclic voltammograms of (a) a microporous thick-film gold electrode and (b) a commercially available planar gold electrode polished to mirror brightness (EG & G) in buffered solutions with (----) and without ( - ) the addition of NADH.
Figure 2 shows the long-term stability of a micro porous gold electrode during the measure ment of NADH.
Figure 3 shows the effect of interfering substances in the case of conventional potentials and potentials made possible by the invention, for the oxidation of NADH.
Figure 4 shows the concentration-current character-istic of a glucose biosensor which, em-ploying the electrode according to the in-vention as a transductor, operates at low polarization voltages.
Figure 1 illustrates the catalytic activity of the electrode:c according to the invention with respect to the anodic oxidation of NADH. In the course of the cyclic voltammogram there appears, in the case of the thick-film gold electrode (a) in the presence of NADH (1), starting at approximately +100 mV (against SCE), a distinct anodic current which is caused by the oxidation of NADH. With the NADH-free buffer (2) this current is absent. With conventional electrodes (b), such behaviour cannot be ob-served under identical experimental conditions (3, 4).
The cyclic voltammograms were recorded in 0.1 M sodium MSE #2105 ~! ~ g 38 1 6 _ ___ phosphate buffer of pH 7.0, which contained 0.1 M sodium perchlorate as the supporting electrolyte. The voltage rate was 10 mV/sec. The NADH concentration was 0 mM (2, 4) and 5 niM (1, 3), respectively.
The thick-film gold electrodes according to the in-vention wE~re incorporated into a flow injection system for detecting 0.5 mM NADH and were operated at +145 mV
(against SCE). Figure 2 shows the change of the anodic 10 current against time for the oxidation of NADH. The ar-rows indicate that in each case freshly prepared NADH so-lutions were used. The drop in the current after the in-troduction of a fresh solution therefore mainly results from the thermal decomposition of NADH in solution (0.1 M
sodium phosphate buffer, pH 7.0 with 0.1 M sodium per-chlorate, room temperature ) and is not due, or only to a limited extent, to electrode fouling.
To clarify the advantages of sensor operation at low polarization voltages, potentially interfering substances (1 and 2 mM hydrogen peroxide, 1 and 2 mM acetaminophen (= paracetamol), serum diluted 1:1 and undiluted) were studied together with NADH (1 and 2 mM) in a flow injec-tion system (Fig. 3). In so doing, in one case, the voltage most common hitherto of (a) +555 mV was applied to the electrodes and (b) in another case at the lower potential of +145 mV (both against SCE), which is suffi-cient to oxidize NADH, using tie electrodes according to the invention. Apart from NADH, nothing was indicated at this low ~aotential except, to a small extent, serum com-ponents.
MSE #2105 X1638 1 6 ~ ._ The glucose biosensor whose calibration curve is re-produced in Figure 4 represents a possible application for the electrodes according to the invention. The en-zyme glucose dehydrogenase was immobilized onto the gold 5 surface by matrix inclusion into a polyvinyl acetate) membrane. The calibration curve was recorded in a flow injection system which was operated with 0.1 M sodium phosphate buffer of pH 7.0 and whose injection volume was 100 N1 at a flow rate of 0.6 ml/min.
MSE #2105
Claims (15)
1. Electrochemical enzyme biosensor which comprises noble metal electrodes and with which an electrochemical oxidation of pyridine nucleotides takes place, characterized in that the noble metal electrodes have a microtexture rich in surface pores and having catalytic properties, as a result of which the overpotential required. for electrochemical oxidation of the pyridine nucleotides is reduced.
2. Electrochemical enzyme biosensor according to Claim 1, characterized in that, owing to the microtextures, the surface area of an electrode is from 3 to 10 times larger than its geometric surface area.
3. Electrochemical enzyme biosensor according to Claim 2, characterized in that the electrode surface comprises noble metal particles having a mean radius of curvature of from 0.05 µm to 4 µm.
4. Electrochemical enzyme biosensor according to Claims 1-3, characterized in that the overpotential required for the anodic oxidation of NADH to NAD+ is reduced to 0-200 mV against a KCl-saturated aqueous calomel reference electrode.
5. Electrochemical enzyme biosensor according to Claims 1-4, characterized in that the electrodes are made of gold.
6. Electrochemical enzyme biosensor according to Claims 1-5, characterized in that the electrodes are coated with a membrane, in which selectivity-determining biocomponents are immobilized.
7. Electrochemical enzyme biosensor according to Claims 1-6, characterized in that the electrodes have been modified by electropolymerization of pyrrole and methylene blue.
8. Electrochemical enzyme biosensor according to Claims 1-7, characterized in that enzymes or other biocomponents are immobilized directly on the electrodes via reactive pendent groups.
9. Method for fabricating noble metal electrodes for an electrochemical biosensor according to Claims 1-6, characterized in that a chemically inert base is coated, by means of screen printing, with a paste comprising a gold powder and a polymeric binder, and the coated base is passed through a furnace having incrementally gradated firing temperature zones, the temperature of the first firing zone being from 300°C to 400°C and that of the last firing zone being from 800°C to 1100°C.
10, Method according to Claim 9, characterized in that the residence time, each time the coated base is passed through a firing zone of the furnace, is from 3 min to 8 min.
11. Electrochemical enzyme biosensor of Claims 1 to 8 wherein the pyridine nucleotide is NADH.
12. Electrochemical enzyme biosensor according to Claim 6 wherein the selectivity-determining biocomponents are enzymes.
13. Electrochemical enzyme biosensor according to Claim 2 wherein the surface area is from 4 to 7 times larger than the geometric surface area.
14. Electrochemical enzyme biosensor according to Claim 3 wherein the mean radius of curvature of the noble metal particles is from 0.5 to 2 µm.
15. Electrochemical enzyme biosensor according to Claim 4 wherein the overpotential required for the anodic oxidation of NADH to NAD+ is reduced to 100 to 150 mV.
Applications Claiming Priority (2)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| DEP4442253.9 | 1994-11-28 | ||
| DE4442253A DE4442253A1 (en) | 1994-11-28 | 1994-11-28 | Electrochemical enzyme biosensor |
Publications (2)
| Publication Number | Publication Date |
|---|---|
| CA2163816A1 CA2163816A1 (en) | 1996-05-29 |
| CA2163816C true CA2163816C (en) | 2005-06-14 |
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Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA002163816A Expired - Fee Related CA2163816C (en) | 1994-11-28 | 1995-11-27 | Electrochemical enzyme biosensor |
Country Status (5)
| Country | Link |
|---|---|
| US (1) | US5783056A (en) |
| EP (1) | EP0714985B1 (en) |
| JP (1) | JPH08220055A (en) |
| CA (1) | CA2163816C (en) |
| DE (2) | DE4442253A1 (en) |
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| US5620850A (en) | 1994-09-26 | 1997-04-15 | President And Fellows Of Harvard College | Molecular recognition at surfaces derivatized with self-assembled monolayers |
| DE19631530C2 (en) * | 1996-07-23 | 2000-03-09 | Inst Chemo Biosensorik | Ion selective sensor |
| US7014992B1 (en) | 1996-11-05 | 2006-03-21 | Clinical Micro Sensors, Inc. | Conductive oligomers attached to electrodes and nucleoside analogs |
| US5866353A (en) * | 1996-12-09 | 1999-02-02 | Bayer Corporation | Electro chemical biosensor containing diazacyanine mediator for co-enzyme regeneration |
| DE19653436C1 (en) * | 1996-12-20 | 1998-08-13 | Inst Chemo Biosensorik | Electrochemical sensor |
| US7560237B2 (en) | 1997-06-12 | 2009-07-14 | Osmetech Technology Inc. | Electronics method for the detection of analytes |
| IL124903A0 (en) * | 1998-06-15 | 1999-01-26 | Bauer Alan Josef | An enzyme biosensor |
| US6322963B1 (en) | 1998-06-15 | 2001-11-27 | Biosensor Systems Design., Inc. | Sensor for analyte detection |
| US6281006B1 (en) * | 1998-08-24 | 2001-08-28 | Therasense, Inc. | Electrochemical affinity assay |
| US6503701B1 (en) | 1999-06-15 | 2003-01-07 | Biosensor Systems Design, Inc. | Analytic sensor apparatus and method |
| US6342347B1 (en) | 1999-10-22 | 2002-01-29 | Biosensor Systems Design., Inc. | Electromagnetic sensor |
| GB0000721D0 (en) * | 2000-01-14 | 2000-03-08 | Univ Wales Aberystwyth The | Protective coatings for metallic electrodes |
| US6824669B1 (en) | 2000-02-17 | 2004-11-30 | Motorola, Inc. | Protein and peptide sensors using electrical detection methods |
| CN1227525C (en) * | 2001-06-14 | 2005-11-16 | 松下电器产业株式会社 | Biosensor |
| EP1523551A4 (en) * | 2001-11-09 | 2007-10-31 | Rosestreet Labs Llc | An enzyme-based system and sensor for measuring acetone |
| US7794994B2 (en) * | 2001-11-09 | 2010-09-14 | Kemeta, Llc | Enzyme-based system and sensor for measuring acetone |
| GB0204232D0 (en) * | 2002-02-22 | 2002-04-10 | Isis Innovation | Assay |
| US20050186652A1 (en) * | 2002-02-22 | 2005-08-25 | Wong Luet L. | Electrochemical detection |
| US7175897B2 (en) * | 2002-12-17 | 2007-02-13 | Avery Dennison Corporation | Adhesive articles which contain at least one hydrophilic or hydrophobic layer, method for making and uses for same |
| US20040197821A1 (en) * | 2003-04-04 | 2004-10-07 | Bauer Alan Joseph | Rapid-detection biosensor |
| EP1751533A2 (en) * | 2004-05-14 | 2007-02-14 | Bayer Healthcare, LLC | Voltammetric systems for assaying biological analytes |
| JP4849434B2 (en) * | 2005-03-31 | 2012-01-11 | 福岡県 | Metal surface modified ceramic scaffolds and their applications |
| JP2006302817A (en) * | 2005-04-25 | 2006-11-02 | Cy Tekku:Kk | Photosynthetic fuel cell using water as material |
| US8008037B2 (en) | 2008-03-27 | 2011-08-30 | Roche Diagnostics Operations, Inc. | Matrix composition with alkylphenazine quaternary salt and a nitrosoaniline |
| JP5930810B2 (en) * | 2011-04-26 | 2016-06-08 | アークレイ株式会社 | Analysis tool |
| DK3423591T3 (en) * | 2016-03-04 | 2024-01-29 | Abbott Diabetes Care Inc | NAD(P)-dependent sensitive enzymes, electrodes and sensors and methods for their production and use |
| FR3063148B1 (en) * | 2017-02-21 | 2022-05-06 | Univ Bordeaux | DEVICE AND ELECTROCHEMICAL METHOD FOR THE MEASUREMENT OF DIFFERENT NON-COMPLEXED FORMS OF SULFUR DIOXIDE IN AN AQUEOUS LIQUID MEDIUM |
| US20220304599A1 (en) * | 2019-08-31 | 2022-09-29 | PercuSense, Inc. | Analyte sensor |
| WO2021221752A2 (en) | 2020-02-06 | 2021-11-04 | Trustees Of Boston University | High throughput assay for identifying microbial redox enzymes |
| US11331020B2 (en) | 2020-02-06 | 2022-05-17 | Trustees Of Boston University | Enzyme-based electrochemical nicotine biosensor |
Family Cites Families (6)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| GB8710472D0 (en) * | 1987-05-01 | 1987-06-03 | Cambridge Life Sciences | Amperometric method |
| US5540828A (en) * | 1987-06-08 | 1996-07-30 | Yacynych; Alexander | Method for making electrochemical sensors and biosensors having a polymer modified surface |
| DE4027728A1 (en) * | 1990-08-31 | 1992-03-05 | Bayer Ag | IMMOBILIZATION OF ORGANIC MACROMOLECULES OR BIOPOLYMERS IN A POLYMER MEMBRANE |
| US5240571A (en) * | 1991-04-24 | 1993-08-31 | University Of Cincinnati | Quantitative method of detection of analytes in aqueous fluids by detection of NADH and NADPH |
| US5217594A (en) * | 1992-01-15 | 1993-06-08 | Enzyme Technology Research Group, Inc. | Convenient determination of trace lead in whole blood and other fluids |
| US5468366A (en) * | 1992-01-15 | 1995-11-21 | Andcare, Inc. | Colloidal-gold electrosensor measuring device |
-
1994
- 1994-11-28 DE DE4442253A patent/DE4442253A1/en not_active Withdrawn
-
1995
- 1995-11-15 DE DE59508908T patent/DE59508908D1/en not_active Expired - Lifetime
- 1995-11-15 EP EP95117976A patent/EP0714985B1/en not_active Expired - Lifetime
- 1995-11-20 US US08/560,939 patent/US5783056A/en not_active Expired - Lifetime
- 1995-11-27 CA CA002163816A patent/CA2163816C/en not_active Expired - Fee Related
- 1995-11-27 JP JP7307233A patent/JPH08220055A/en active Pending
Also Published As
| Publication number | Publication date |
|---|---|
| US5783056A (en) | 1998-07-21 |
| EP0714985B1 (en) | 2000-12-20 |
| DE4442253A1 (en) | 1996-05-30 |
| DE59508908D1 (en) | 2001-01-25 |
| JPH08220055A (en) | 1996-08-30 |
| CA2163816A1 (en) | 1996-05-29 |
| EP0714985A1 (en) | 1996-06-05 |
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