Disposable electrochemical bio-sensor for the quantitative determination of analyte concentrations in fluids

ELECTROCHEMICAL ONE WAY BIO SENSOR FOR THE QUANTITATIVE DETERMINATION OF ANALYTICAL CONCENTRATIONS IN LIQUIDS

DESCRIPTION

The present invention relates to a disposable biosensor comprising a carrier material, electrical conductor tracks, an electrode system comprising a counter electrode and a working electrode formed from a reaction layer, a dielectric insulator layer which covers the carrier material and the electrical conductor tracks and has cutouts for contacting a potentiostat unit and for the electrode system , as well as an analyzer-recognizing biocomponent. The invention also relates to methods for determining analytes in a liquid sample using the disposable biosensors according to the invention, the use of slightly subliming compounds as electron transfer mediators of an electrochemical sensor for transferring electrons from an enzyme to an electron-conducting material, and the use of the disposable biosensors according to the invention for determining Analyte concentrations in body or sample fluids.

Electrochemical detection systems based on mediator-modified biosensors for the specific quantitative determination of an analyte in a liquid sample have already been described in detail in the literature and in many patents (for example in EP 0 552 223, US 5,288,636, US 4,711, 245). These are mostly biosensors for analytes that can be converted specifically by enzymes of the oxidoreductase class (lactate, cholesterol, ethanol, hydrogen peroxide, uric acid, etc.). Some mediator-modified biosensors are already available as commercial products. The best-known example of such a biosensor is the glucose biosensor, which is used in medical diagnostics by diabetics to determine glucose concentrations in whole blood. The detection principle of a mediator-modified biosensor is described for example in EP 0 552 223 for a glucose biosensor. The enzyme used in a glucose biosensor is glucose oxidase (GOD). The GOD has a prosthetic group in the form of a flavin adenine dinucleotide molecule (FAD) and catalyzes the reaction of glucose to gluconolactone according to equation (I):

(I) glucose + GOD-FAD => gluconolactone + GOD-FADH₂

In principle, it is now possible to use the reduced enzyme (GOD-FADH₂) with oxygen (0₂) to GOD-FAD with formation of hydrogen peroxide (H₂0₂) to oxidize according to equation (Ila) and then to determine the glucose concentration in the sample directly by determining the hydrogen peroxide concentration using a platinum electrode.

(lla) GOD-FADH₂ + O₂ o GOD-FAD + H₂0₂

However, in most versions of the glucose biosensor, an electron transfer mediator is used to quantify the amount of glucose in the sample. In a second reaction step, this mediator oxidizes the enzyme, so that it is available again for reaction (I):

(llb) GOD-FADH2 + MEDox = GOD-FAD + MED_rΘd

The glucose concentration in the sample can now be determined by using the reduced mediator (MED_red) oxidized again at the working electrode of the biosensor and the current flowing was quantified:

(III) MED_red = MEDox + e

In order to allow this reaction to take place, a potentiostat unit is required which specifies a suitable measuring potential and measures the flowing current. The current is proportional to the concentration of glucose in the sample. The principle of detection of a mediator-modified biosensor has the advantage over the direct quantification of hydrogen peroxide on a platinum electrode that it is possible to work with a low working potential (0-350 mV vs. Ag / AgCI). This serves to reduce, ideally even to avoid interfering interferences that can arise from the oxidation of accompanying substances (interfering substances in the blood are, for example, ascorbic acid, uric acid or paracetamol).

Another way to use modified biosensors is to quantify analytes that can be specifically integrated in an enzyme-labeled assay. For this purpose, the analyte is specifically bound to a capture molecule (analyte-recognizing biocomponent) immobilized on the surface of the sensor. In turn, a molecule labeled with an enzyme is coupled to this complex of capture molecule and analyte. The concentration of the bound analyte can be determined by quantifying the amount of enzyme immobilized on the surface of the biosensor via the assay. This is done by adding a defined concentration of a substrate, which is implemented by the enzyme. The mediator of the biosensor regenerates the enzyme bound to the assay according to equation (IIb). The detection reaction is again the oxidation or Reduction of the mediator on the biosensor surface (see equation III). The current to be measured is proportional to the amount of enzyme labeled and thus also proportional to the amount of analyte bound in the assay. This sensor principle is used, for example, in immunosensorics or DNA analysis.

Compounds that are used as mediators in the biosensors should be distinguished by the fact that they have a low redox potential, that they achieve high conversion rates with the enzyme, that they are not readily soluble in the reduced state and that they are stable and inert to other accompanying substances are. Many compounds are published in the literature and in numerous patents that can be used as mediators in biosensors. A large group of mediators are organometallic compounds. Examples of these are ferrocene derivatives (British Patent 8132034, US 4,711, 245), transition metal complexes (US 5,710,011), e.g. B. Potassium hexacyanoferrate (US 5,645,709), nickel or cobalt bipyridyl and osmium-containing organic compounds (US 5,846,702). However, these organometallic mediators have the disadvantage that they have a comparatively low conversion rate and accordingly have to be used in high concentrations. This is particularly disadvantageous for the use of these mediators in disposable biosensors in the form of test strips. Another disadvantage of organometallic mediators is their good solubility in water. The consequence of this is that the mediators can diffuse out of the electrode space relatively easily when they come into contact with the sample liquid, so that the sensitivity of the biosensor deteriorates due to a decrease in the mediator concentration in the electrode space and it also due to the contact of the mediators with accompanying substances , which are present especially when analyzing blood samples, can lead to disruptive side reactions. Compounds which have a lower solubility and which also interact to a lesser extent with accompanying substances are therefore more advantageous.

These desired, advantageous properties are generally fulfilled by molecules which form quinoid structures (hydroquinone / benzoquinone, p-aminophenol / p-iminoquinone) and which are therefore also used as mediators in biosensors. Frequently used compounds are also tetrathiafulvalene (TTF), tetracyanoquinodimethane (TCNQ) or N-methylphenazinium (NMP) (US 5,876,952). Many of these compounds are particularly good electron donors or acceptors due to their photosensor properties and are therefore often used as mediators in biosensors. In this context, for example, N, N, N \ N should be mentioned^'-tetrakis- (2nd^'hydroxyethyl) -p-phenylenediamine (THEPD) and N, N, N^', N^'-tetrakiscarboxymethyl-p-phenylenediamine (TCPD) (US 5,609,749) and 1,4-diamino-2,3,5,6-tetramethylbenzene, 2,5-diethyl-1,4-bis (dimethylaminobenzen), N- (4- Morpholinophenylhexahydroazepine),

Meldola blue and others (WO 9207263). These molecules are essentially derivatives of phenylenediamine (PPD) which, like the N, N, N^', N'-tetraphenylenediamine (TMPD), also has good donor / acceptor properties (J. Anal. Chem. USSR 45 (1990)). Compared to the organometallic mediators, the derivatives of the PPD are characterized by an increased conversion rate and a lower oil quality. The disadvantage of these slightly subliming compounds is that they have a relatively high vapor pressure. This leads to the fact that the mediator, after being applied to the substrate of the biosensor, evaporates over time and thus the sensitivity of the biosensor decreases with prolonged storage. Another problem is the fact that the derivatives of the PPD diffuse out of the electrode space when in contact with the aqueous sample liquid due to a low solubility in water, even if compared to the organometallic mediators, and consequently a lower sensitivity of the biosensor is observed due to the decrease in the mediator concentration becomes.

The object underlying the present application was therefore the development of a mediator-modified biosensor which, despite the use of slightly subliming compounds as mediators, does not have the above-mentioned disadvantages of a loss of sensitivity due to prolonged storage or as a result of contact with the sample liquid.

The object was achieved according to the invention in that the electrode system was provided with a polymeric protective layer in a biosensor of the type mentioned at the outset. This protective layer not only increases the sensitivity of the biosensor by keeping the mediator that is detached from the electrode near the electrode, but above all achieves a significant improvement in the long-term stability of the biosensor despite the use of slightly subliming mediators. This is mainly due to the fact that the polymeric protective layer counteracts the evaporation of the mediator. The linear measuring range of the biosensor can also be extended by the effect of this polymeric protective layer as a diffusion barrier.

Many biosensors presented in publications and patents have been provided with polymer layers for a variety of reasons. For example, polymeric membranes are used to immobilize the enzymes on the sensor surface (EP 0 851 244, EP 0 894 869, EP 0 856 586, EP 0 909 952). In other applications, a polymer is used to build up a so-called reaction layer over the working electrode, which contains both the mediator and the Enzvm (patents of Matsushita Electric Ind. Co. LTD. JP. For example: US 5,192,415, EP 0 901 018, WO 9835225 ). Polymer layers can also be used to prevent interference from interfering substances or to protect the sensor surface from contamination (the latter is particularly important for reusable biosensors and the development of in vivo sensors). Nafion layers are often used to reduce interference from anionic interfering substances (e.g. ascorbate) (US 5,312,590). A layer of polyacrylamide can serve to reduce the influence of the hematocrit of a blood sample on the sensor signal (EP 0 769 558). Ultimately, polymeric layers can also be used to create a passivation layer that prevents non-specific bonds to the sensor electrode (DE 198 06 642). However, the use of a polymeric protective layer to prevent the evaporation of a slightly subliming mediator and thus also to improve the long-term stability of biosensors has not previously been described in the literature.

The subject matter of the present invention relates to a disposable biosensor comprising a carrier material, electrical conductor tracks, an electrode system comprising a counter electrode and a working electrode formed from a reaction layer, a dielectric insulator layer which covers the carrier material and the electrical conductor tracks and cutouts for the contacting of a potentiostat unit and for that Has electrode system, and an analyzer-recognizing biocomponent. The disposable biosensor according to the invention is characterized in that the reaction layer contains a slightly subliming mediator and that the electrode system is covered by a polymeric protective layer. The disposable biosensor according to the invention is produced in the form of a chip and used in combination with a potentiostat unit. It is used for quick on-site quantification of analyte concentrations in liquids. The detectable analytes are substances that can be converted specifically by oxidoreductases or that can be integrated in an enzyme-labeled immuno or receptor assay. The quantification is carried out amperometrically or cyclovoltametric.

The structure of one of the preferred embodiments of the biosensor according to the invention is shown as an example in FIG. 1. The biosensor comprises a carrier material (1) on which electrical conductor tracks (2), an insulator layer (3), an electrode system comprising a counter electrode (6) and a working electrode (5), and a polymeric protective layer (8) covering the electrode system are preferably arranged have been applied by mask or screen printing, and an analyzer-recognizing biocomponent. The working electrode (5) is formed from a reaction layer containing a slightly subliming mediator and an electron-conducting material. The conductor tracks serve to connect the respective electrodes to the connection points (4) via which the biosensor is contacted with the potentiostat unit. The insulator layer (3) covers the conductor tracks and has cutouts for the connection points (4) and for the electrode system and serves to prevent parasitic currents. In the preferred embodiment of the disposable biosensor according to the invention shown in FIG. 1, in addition to the counter electrode (6) and the working electrode (5), the electrode system also comprises a reference electrode (7). By using this reference electrode, a significant improvement in measuring accuracy can be achieved.

The conductor tracks (2) consist of a polymer paste containing carbon, silver, platinum or goid particles as essential components. The conductor tracks (2) are preferably applied to the carrier material (1) by screen or mask printing and then cured.

The carrier material (1) used to manufacture the disposable biosensors preferably consists of polycarbonate, polyvinyl chloride (PVC), polyethylene terephthalate (PET), polypropylene, polyester or polystyrene.

To produce the counter electrode (6) and the reference electrode (7), a conductive paste containing silver particles and silver chloride, to which a readily soluble chloride salt (e.g. sodium chloride or potassium chloride) has been added, is preferably used. The counter electrode (6) and the reference electrode (7) are preferably applied to the carrier material (1) by screen or mask printing and then cured. The reaction layer from which the working electrode (5) has been formed contains, in addition to the slightly subliming electron transfer mediator, an electron-conducting material which consists of a curable polymer paste containing carbon, platinum, palladium, rhodium or gold particles as essential components. The reaction layer is preferably applied to the carrier material (1) by screen or mask printing and then cured.

The lightly subliming mediators used in the present invention are compounds that have a vapor pressure of at least MO^"5 mmHg at 25 ° C. The mediators preferably have the structure shown in Figure 2, in which the residues R, R², R³ and R⁴ may be the same or different and represent a hydrogen atom, a CrCio-alkyl radical, preferably a Ci-Cs-alkyl radical, or an aryl radical. Preferred mediators are p-aminodiphenylamine, p-phenylenediamine or N, N, N^', N^'-Tetramethyl-p-phenylenediamine.

The complete electrode system is covered by a hydrophilic protective layer (8) stabilizing the mediator in the form of a polymeric hydrogel, which counteracts the evaporation of the electron transfer mediator and at the same time increases the sensitivity of the biosensor and acts as a diffusion barrier. The preferably screen or mask-printed polymeric protective layer (8) essentially consists of a hydrogel made of polyvinylpyrrolidone (PVP), polyethylene oxide (PEO), starch or gelatin.

An enzyme, a DNA molecule, a capture antibody or a receptor molecule can function as the analyte-recognizing biocomponent. An oxidoreductase, a dehydrogenase, a β-galactosidase or an alkaline phosphatase is preferably used as the enzyme. Particularly preferred enzymes are glucose oxidase (GOD), lactate oxidase (LOD), uricase, ascorbate oxidase, ethanol oxidase, cholesterol oxidase or peroxidase (POD).

In the preferred embodiment of the biosensor according to the invention shown in FIG. 1, a plastic fabric (9) with a one-sided adhesive tape (10) is also attached above the electrode system of the biosensor. The one-sided adhesive tape (10) has a recess into which the liquid sample can be brought, so that the sample contacts the electrodes via the plastic fabric (9), which electrodes are supplied with a measuring voltage by the potentiostat unit. However, this tissue is used not only to quickly and evenly distribute the sample over the electrodes, but also to protect the biosensor surface from destruction or contamination, which can occur, for example, when capillary blood is applied as a sample with the finger. The plastic fabric (9) preferably consists of a polyethylene, polypropylene or polyamide network, while the tape (10) consists of a one-sided adhesive polyester, PVC or paper adhesive film.

Depending on the location of the analyzer-recognizing biocomponent and the structure of the reaction layer, five different embodiments of the biosensor according to the invention can be distinguished.

1. In a first embodiment of the biosensor according to the invention, the reaction layer consists of a single layer containing the electron-conducting polymer paste, the slightly subliming mediator and the analyte-recognizing biocomponent.

2nd In a second embodiment of the biosensor according to the invention, the reaction layer consists of a single layer containing the electron-conducting polymer paste and the slightly subliming mediator. In this embodiment, however, the analyte-recognizing biocomponent is located in the polymer protective layer (8).

3rd In a third embodiment of the biosensor according to the invention, the reaction layer consists of two layers, the first layer being formed from the electron-conducting polymer paste and the slightly subliming mediator and being covered by a second, hydrophilic polymeric top layer which acts as a carrier of the analyte-recognizing biocomponent. 4th In a fourth embodiment of the biosensor according to the invention, the reaction layer also consists of two layers, in which case the first layer is formed from the electron-conducting polymer paste and is covered by a second, hydrophilic polymer cover layer, which carries the analyte-recognizing biocomponent and the slightly subliming mediator acts.

5. In a fifth embodiment of the biosensor according to the invention, the reaction layer likewise consists of two layers, the first layer being formed from the electron-conducting polymer paste and being covered by a second, hydrophilic polymer cover layer which acts as a carrier of the slightly subliming mediator. In this embodiment, the analyte-recognizing biocomponent is located in the polymer protective layer (8).

The analyte-recognizing biocomponent can be immobilized either directly or via an avidin / biotin interaction (see Figures 18 and 20) in the reaction layer or in the polymer protective layer (8).

The polymeric top layer optionally contained in the reaction layer essentially consists of a hydrogel of polyvinylpyrrolidone (PVP), polyethylene oxide (PEO), starch or gelatin, which has preferably been printed on by screen or mask printing.

The present invention also relates to a method for determining analyte concentrations in aqueous solutions using the disposable biosensors according to the invention.

When using that embodiment of the biosensor in which the analyte-recognizing biocomponent is an enzyme, e.g. B. is glucose oxidase, lactate oxidase, alcohol oxidase, peroxidase or uricase, a liquid sample is first placed on the electrode system of the biosensor. The analyte to be determined, e.g. B. glucose, is then by the respective enzyme, as described for the case of Glucose Mar.h fiisfis in Gleichunα (\). reduced. The reduced Enzvm then reduces the mediator in a further reaction step according to equation (IIb). The mediator is then oxidized and the current flowing, which is proportional to the analyte concentration, is determined via a potentiostat unit, which specifies a suitable measurement potential.

If the biosensor according to the invention is used as an enzyme-labeled immunoassay, as a DNA assay or as a ligand assay, then in this case too, a liquid sample is first applied to the electrode system of the biosensor. The antigen contained in the sample liquid, the sequence-specific DNA molecule or the ligand then bind to the analyte-recognizing biocomponent (antibody, DNA probe or receptor) immobilized in the reaction layer or the polymer protective layer (8). The sensor surface is then washed and then covered with a solution which contains an analyte-specific enzyme conjugate which recognizes the sequence-specific DNA, the antigen or the ligand. This enzyme conjugate binds to the analyte (antibody, DNA probe or receptor) and the amount of the enzyme conjugate bound in the reaction layer or the polymeric protective layer (8) is proportional to the amount of analyte in the sample liquid initially applied. After thorough washing of the surface of the working electrode again, it is coated with a further solution containing a substrate of the enzyme conjugate. The substrate is converted by the enzyme and the reduced or oxidized enzyme then reacts with the mediator in accordance with equation IIb). The current flowing is then determined via a potentiostat unit, which specifies a suitable measurement potential.

The present invention also relates to the use of slightly subliming compounds with a vapor pressure of at least 1 -10^"5 mmHg at 25 ° C as electron transfer mediators for the transfer of electrons from an enzyme to an electron-conducting material, these mediators being located in a reaction layer which is covered by a polymeric protective layer (8). The polymeric protective layer (8) counteracts the sublimation of the mediator, so that the long-term stability of the sensor can be extended considerably. The mediators that are preferably used to transfer electrons from an enzyme to an electron-conducting material are slightly subliming compounds with the structure (I),

<img file="WO0202796A2_D0001.tif" />

in which the residues R¹, R², R³ and R⁴ can be the same or different and a hydrogen atom, a -C-Cιo-alkyl radical, preferably a C-_t-Cs alkyl radical, or an aryl radical. Preferred mediators used to transfer electrons from an enzyme to an electron-conducting material are p-aminodiphenylamine, p-phenylenediamine or N, N, N \ N^'-Tetramethyl-p-phenylenediamine.

The present invention further relates to the use of the disposable biosensor according to the invention for determining analyte concentrations in body or sample fluids. The analytes detectable with the biosensor according to the invention are preferably glucose, lactate, uric acid, ascorbate, ethanol, cholesterol or hydrogen peroxide. Sequence-specific DNA fragments, specific antigens or ligands can also be detected with special embodiments of the biosensor according to the invention. The liquids that can be analyzed with the present biosensor are aqueous synthetic solutions, nutrient solutions, or also blood, blood serum, blood plasma or urine. EXAMPLES

General information on the characterization of the sensors

The redox behavior of the mediator used in the disposable biosensor can be determined using cyclic voltammograms in a buffered solution (PBS, pH 7.4, 154 mM CI^") represent. Peaks occurring therein illustrate the electrochemical oxidizability or reducibility of the mediator. The peak potentials determine the working potential of the biosensor. The peak currents are a measure of the oxidized or reduced amount of substance. Falling peak currents after longer biosensor storage times indicate the loss or decay of the mediator.

To check the stability of a mediator, cyclic voltammograms in buffered aqueous solutions are carried out at regular intervals with unused biosensors and the peak currents obtained are compared.

Both cyclic voltammograms and amperometric measurements in analyte-containing buffer show the ability of the respective mediator to regenerate the enzyme and transfer the electrons to the electrode. This manifests itself in both experiment forms in an increase in the currents compared to measurements which are carried out in analyte-free buffer. The sensitivity of the biosensors can be determined using amperometric measurements at different analyte concentrations. This is defined as the increase in current with increasing concentration. Both the change in sensitivity that occurs with increasing sensor age and the change in peak currents in the cyclic voltammogram serve to characterize the long-term stability.

A) Biosensors for determining glucose concentrations in sample liquids

With the sensor structure described above, differently designed iiiknς .hinspnsnrAn can be produced. The different modifications of these glucose biosensors differ, inter alia, in the choice of the compound which is used as a mediator or in the material of the polymeric protective layer. In the listed examples, 1-3 glucose oxidase (from Aspergillus niger) is used as the enzyme. The protective and sensitivity-increasing effect of the polymer protective layer is illustrated by comparing measurement results from biosensors with and without a protective layer.

Example 1 Electrode material: carbon paste

Mediator: p-aminodiphenylamine (ADPA)

Enzyme: glucose oxidase (Aspergillus niger)

Protective layer: polyvinyl pyrrolidone (PVP)

Embodiment The reaction layer consists of a single layer containing an electron-conducting material, the mediator and the analyte-recognizing biocomponent (glucose oxidase)

The peak current (oxidation current) of an ADPA-modified biosensor without a protective layer is 216 nA at 200 mV (vs. Ag / AgCI) at the beginning of the series of measurements (Figure 3). The peak current can be increased to 2000 nA by the signal-increasing effect of a PVP protective layer (Figure 4). While the ADPA biosensors without a protective layer lose 55% of their oxidizable ADPA amount within 54 days, this amount remains constant during this period for biosensors with a PVP protective layer.

The cyclic voltammograms in buffered glucose solution <img file="WO0202796A2_D0002.tif" /> mM) show that ADPA is suitable as a mediator. The current in the potential range of the oxidation of ADPA increases to 550 nA due to the presence of glucose in biosensors without a protective layer and to approximately 6000 nA in biosensors with a protective layer. While the ADPA biosensors without a protective layer show significant signal losses within 54 days (Figure 5), the signal value of the PVP-covered biosensors remains constant during this period (Figure 6). The use of the PVP protective layer in the ADPA biosensors also brings about a significant improvement with regard to the sensitivity of the sensors. Initially, the sensitivity of the unprotected ADPA biosensors was 204 nA / mM (Figure 7). With protected biosensors, the sensitivity is initially 325 nA / mM (Figure 8). While the PVP protective layer ensures that this sensitivity of 325 nA / mM is retained even after 54 days, it drops to 136 nA / mM within 54 days for the unprotected biosensors. The extension of the linear measuring range of the ADPA biosensor by the PVP protective layer is also striking. Without this protective layer, saturation of the sensor signal that is already occurring with 4 mM glucose is observed, whereas with protected biosensors this concentration is still detected in the linear range of the biosensor.

Example 2

Electrode material: carbon paste

Mediator: N, N, N^', N^'-Tetramethyl-phenylenediamine (TMPD) enzyme: glucose oxidase (Aspergillus niger)

Protective layer: polyvinyl pyrrolidone (PVP)

Embodiment The reaction layer consists of a single layer containing an electron-conducting material, the mediator and the analyte-recognizing biocomponent (glucose oxidase)

TMPD-modified biosensors without a protective layer show an oxidation peak with a peak current of 90 nA (vs. Ag / AgCI) in a cyclic voltammogram in giose-free buffer at 100 mV (Figure 9). This signal can be amplified to 850 nA by means of a PVP protective layer (Figure 10). The oxidizable amount of unprotected biosensors is reduced by 57% within 46 days. In the case of TMPD biosensors with a PVP protective layer, on the other hand, the oxidation current and thus the amount of TMPD that can be oxidized remains almost constant during this period.

TMPD shows in the cyclic voltammograms in buffered glucose solution <img file="WO0202796A2_D0003.tif" /> mM) the usability as a mediator. The current in the potential range for the oxidation of TMPD is initially 450 nA for biosensors without a protective layer (Figure 11) and 3500 nA for biosensors with a protective layer (Figure 12). This value can also be kept almost constant only with biosensors with a PVP protective layer. Without a protective layer, the oxidation current drops to 160 nA within 46 days. The stabilizing and signal-increasing effect of the PVP protective layer is also evident when considering the sensitivity. This increases due to the PVP protective layer from 60 nA / mM (Figure 13) to 341 nA / mM (Figure 14) and can be kept constant at this level for 46 days. As can already be observed with the ADPA biosensor, the PVP protective layer also extends the linear measuring range with the TMPD biosensor, so that a glucose concentration of 4 mM can still be detected.

Example 3

Electrode material: carbon paste

Mediator: N, N, N^', N^'-Tetramethyl-phenylenediamine (TMPD)

Enzyme: glucose oxidase (Aspergillus niger) Protective layer: starch

Embodiment The reaction layer consists of a single layer containing an electron-conducting material, the mediator and the analyte-recognizing biocomponent (glucose oxidase)

A TMPD biosensor can also be stabilized with a protective layer of starch. This layer increases the signal intensity in the cyclic voltammogram in buffered glucose solution to 175 nA (Figure 15) (compared to 90 nA without a protective layer, Figure 9) and remains almost constant over a period of 46 days. In a buffered glucose solution<img file="WO0202796A2_D0004.tif" /> mM), oxidation currents between 470 and 520 nA (Figure 16) are obtained over a period of 46 days compared to 450 nA without a protective layer (Figure 11). The signal-increasing effect of starch is therefore not as intense as with PVP, but stabilization can still be achieved with this protective layer. This is also evident in the sensitivity studies (Figure 17). B) Biosensors for determining the concentrations of sequence-specific DNA single strand fragments (Example 4) or of anti-quenes in sample liquids (Example 5)

In addition to analytes that are converted by specific enzymes, the sensor structure described above can also be used to quantify single-stranded DNA fragments or even antigens in sample or body fluids.

Example 4

Electrode material: carbon paste Mediator: p-aminodiphenylamine (ADPA)

Enzyme: peroxidase

Biocomponent: biotinylated single-stranded DNA protective layer: starch

Embodiment: The reaction layer consists of a single layer containing an electron-conducting material, the mediator and the analyzer-recognizing biocomponent (DNA probe).

The biotinylated sequence-specific DNA probe is immobilized on the reaction layer via avidin (Figure 18). The amount of complementary DNA strand to be determined is marked with digoxigenin. After hybridization of this complementary DNA strand to the immobilized DNA single strand, detection is carried out after adding an anti-digoxigenin-antibody conjugate (enzyme conjugate). The peroxidase catalyzes the reduction of hydrogen peroxide (substrate of the enzyme conjugate) to water, whereby the peroxidase is oxidized. The peroxidase is regenerated by the mediator ADPA, the oxidized mediator is then reduced again at the working electrode at a potential of -150 mV (vs. Ag / AgCI). Figure 19 shows a calibration line for the detection of a 20mer oligonucleotide. The detection limit is 80 amol of DNA. Example 5

Electrode material: carbon paste

Mediator: p-aminodiphenylamine (ADPA)

Enzyme: peroxidase

Biocomponent: scavenger antibody

Protective layer: starch

Embodiment: The reaction layer consists of a single layer containing an electron-conducting material, the mediator and the analyzer-recognizing biocomponent (capture antibody).

The capture antibody is either directly immobilized or is fixed on the surface using previously immobilized avidin (Fig. 20). After the analyte from the sample to be examined has been bound to the analyte-specific capture antibody, detection is carried out after addition of a second analyte-specific antibody-peroxidase conjugate (enzyme conjugate). The peroxidase catalyzes the reduction of H₂0₂ to H₂O, which oxidizes the peroxidase. The peroxidase is regenerated by the ADPA mediator; the oxidized mediator is reduced again at the working electrode (at E = -150 mV vs. Ag / AgCI). Figure 21 shows a calibration line using the example of interleukin-4. The detection limit is 200 amol of lnteιieukin-4.