JPH06174679A - Biosensor - Google Patents

Abstract

PURPOSE:To maintain the temperature of an electrode system and a sample liquid when dropping a sample to be constant and to achieve a highly accurate measurement by providing an electrical heating element at the other surface side of an insulation substrate or in an insulation substrate. CONSTITUTION:An electrode system consisting of at least a measurement electrode 3 and a counter electrode 2 near it is provided on an insulation substrate 1. An electrical heating element 10 is provided in the substrate 1 or on the substrate 1 of the reverse surface of the electrode system. The element 10 is provided with a function for transferring the thermal energy to the electrode system and the sample system rapidly and then maintaining it to be a constant temperature, thus performing measurement stably without being affected by the external air temperature and the sample liquid temperature. Also, it is proper to maintain a desired temperature by applying a specific potential or conducting a specific current value to a printing resistor. Namely, by applying a specific potential to a terminal part 11 of the resistor 10, the surface temperature of the electrode system is maintained to be constant. Also, the sample is dropped on the electrode system, a specific pulse potential is applied between electrode terminals 8 and 7, and current flowing between both electrodes 2 and 3 is measured, thus determining it as response current according to enzyme reaction.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a biosensor, and more particularly to an enzyme sensor suitable for measuring the concentration of a trace amount of a biological substrate contained in body fluid components such as blood and urine.

[0002]

2. Description of the Related Art Analytical methods utilizing the excellent substrate specificity of enzymes are used in fields such as clinical analytical chemistry, food manufacturing, and environmental chemistry. In particular, in the field of clinical analytical chemistry, enzyme sensors capable of selectively detecting biological substrates such as glucose, urea and uric acid have been developed. These enzyme sensors are composed of an electrode such as an oxygen electrode or a platinum electrode and an enzyme-immobilized membrane, and a substance on which the enzyme specifically acts by reading a substance change due to an enzyme reaction as a change amount of an electric signal by the electrode. The concentration of is measured. For example, in a glucose sensor or the like, an electrode active substance such as hydrogen peroxide or oxygen produced or consumed according to the following formula is monitored by an electrode to measure the biological substrate concentration.

[0003]

[Chemical 1]

However, the enzyme sensor based on such a principle has the following problems. Substrate requires stoichiometric oxygen for reaction, but in actual measurement, for example, when measuring glucose concentration in blood of diabetic patients with glucose sensor, the amount of dissolved oxygen in body fluid is insufficient. . Therefore, it is necessary to dilute the sample blood or supplement oxygen with some method. Further, when electrically monitoring hydrogen peroxide, a reducing substance such as ascorbic acid causes measurement errors, and it is necessary to take some measures to remove these errors. Furthermore, since the conventional sensor requires the enzyme-immobilized membrane to be attached to the oxygen electrode or the hydrogen peroxide electrode, there is a limit to miniaturization.

On the other hand, in order to solve these problems, as an enzyme electrode for directly detecting electron transfer accompanying an enzyme reaction,
An enzyme electrode using a conductive polymer and an enzyme electrode using an electron mediator have been proposed. However, an enzyme electrode using a conductive polymer has an advantage that it is not affected by dissolved oxygen, but has problems such as low response and long response time. Furthermore, when the method of capturing the enzyme in the polymerized membrane during the electropolymerization is adopted, it is difficult to control the amount of the immobilized enzyme, and when it is used as an enzyme electrode, it is a substrate over time due to the elimination of the enzyme. A decrease in responsiveness is unavoidable. In addition, even an enzyme electrode using an electron mediator has low conductivity and insufficient response and response time.Because the electron mediator is dispersed in the carbon paste, the elution and desorption of the electron mediator is performed. However, there is a problem that the responsiveness decreases with time.

By the way, in the case of actually quantifying a specific component in a biological sample using such an enzyme electrode, it is not only necessary to perform measurement with high accuracy, but also operations such as dilution and agitation of the sample solution are not required, which is simple. It is desirable to be able to measure quickly and quickly. Further, in the case of a sample such as blood, there are many restrictions on the amount of sample that can be used, and measurement with a small amount of sample is desired.
None of the conventional enzyme electrodes can perform accurate measurement easily and quickly and even with a small amount of sample for a long period of time.

Therefore, as a solution to these drawbacks of conventional enzyme electrodes, the present inventor previously proposed an enzyme electrode having a conductive layer containing an organic charge transfer complex crystal on the surface of a conductive substrate ( Japanese Patent Application No. 2-244840). By directly detecting the electron transfer accompanying the enzyme reaction, this enzyme electrode has the advantages that it is not affected by dissolved oxygen and that it is less affected by interfering substances. It has an advantage that an accurate response can be given. Further, the present inventor has made further improvements in this enzyme electrode, and has proposed an enzyme electrode capable of efficiently performing electron transfer accompanying an enzymatic reaction and having improved responsiveness (Japanese Patent Application No. 3-7908, Japanese Patent Application No. 3-86884). Furthermore, an enzyme electrode containing an organic charge-transfer complex crystal in a conductive layer is used as a measurement electrode, and a miniaturized biosensor is prepared by providing a counter electrode in the vicinity of the measurement electrode, which does not require operations such as dilution and stirring of a sample solution. It enabled measurement with a small amount of sample (Japanese Patent Application No. 4-11346). We also proposed improvement by providing a liquid retaining layer and a filtrate layer on the electrode system including the measuring electrode.
(Japanese Patent Application No. 4 ─ 278828). When such a biosensor is applied to a field such as self-blood glucose level control, it is further desired to eliminate the variation in response due to the temperature of the outside air or the sample liquid and perform stable and highly accurate measurement.

[0008]

DISCLOSURE OF THE INVENTION The present invention provides a biosensor capable of measuring a small amount of sample by using an enzyme electrode containing an organic charge transfer complex, without requiring dilution and stirring of the sample solution. An object of the present invention is to provide a biosensor that can perform accurate measurement without being affected by the temperature of the outside air or the sample solution.

[0009]

[Means for Solving the Problems] In the biosensor in which an enzyme electrode (measuring electrode) containing an organic charge-transfer complex crystal and a counter electrode are arranged on an insulating substrate, the inventors of the present invention have It was found that the temperature of the electrode system and the sample solution at the time of dropping the sample can be kept constant by providing the heating element in the, and stable and accurate measurement can be performed, and the present invention has been completed.

(Summary of the Invention) The present invention provides an electrode system comprising a measuring electrode and a counter electrode on one surface of an insulating substrate, and the measuring electrode comprises an enzyme and an electron in an electrode containing an organic charge transfer complex crystal as a conductive layer. A biosensor in which a mediator is immobilized is characterized in that a heating element is provided on the other surface side of the insulating substrate or in the insulating substrate.

Furthermore, the present invention provides a biosensor in which a filtrate layer and / or a liquid retaining layer is provided on the measuring electrode or an electrode system including the measuring electrode in the above biosensor, and a filtration layer and / or a liquid retaining layer of this biosensor. TECHNICAL FIELD The present invention relates to a biosensor having a layer carrying an anticoagulant. These biosensors may be provided with a compensation electrode on which the enzyme is not immobilized. Further, the biosensor of the present invention is particularly suitable when the enzyme is a redox enzyme.

[0012]

The biosensor of the present invention has a structure in which an electrode system including at least a measurement electrode and a counter electrode provided in the vicinity thereof is provided on an insulating substrate, and a heating element is provided on the insulating substrate. In the biosensor of the present invention, by covering the electrode system with the filtrate layer and / or the liquid-retaining layer, it becomes possible to perform an accurate measurement even with a very small amount of sample, and in at least one of the filtrate layer and the liquid-retaining layer. It is also possible to carry an anticoagulant and carry out a more rapid measurement. In addition, by providing a compensation electrode in which the enzyme is not immobilized, adsorption of proteins in the sample,
It is also possible to perform measurement with higher accuracy by eliminating the influence of side reactions.

The insulating substrate used in the biosensor of the present invention has an electrical insulating property, has no elution of impurities in the substrate with respect to the sample liquid, has an adhesive property with a heating element described later, and has an appropriate heat resistance. As long as it has conductivity, it can be rigid or flexible and is not particularly limited. From such a viewpoint, as the organic material, general-purpose thermoplastic resins such as polyvinyl chloride, polyethylene, polypropylene, polyethylene terephthalate, and acrylic resins, epoxy resins, phenol resins, unsaturated polyester resins, and thermosetting resins such as melamine resins are used. In addition to the resin, it is also possible to use a composite material in which these resins are combined with glass fiber, paper base material, inorganic filler and the like. Further, as the inorganic material, ceramics such as alumina, a glass plate or the like can be used.

In the biosensor of the present invention, in the insulating substrate,
Alternatively, it is characterized in that the heating element is provided on the insulating substrate on the back surface of the electrode system. Enzymatic reactions are generally susceptible to temperature, and in the case of glucose oxidase, for example, the reaction rate increases with increasing temperature up to around 40 ° C. Therefore, when using a biosensor that uses an enzyme reaction for self-blood glucose level control, etc., the electrode reaction is affected by the temperature of the outside air or sample solution at the time of measurement, and as a result, the response varies and it is difficult to obtain a stable response. Sometimes. With the biosensor of the present invention, there are few errors due to the influence of such temperature changes, and stable and highly accurate measurement can be performed.

The heating element has a certain amount of heat generation function in response to the application of external electrical, optical or mechanical energy, and the heat energy is rapidly transmitted to the electrode system and the sample solution to maintain a constant temperature. There is no particular limitation as long as it is an element having the function of performing. From the viewpoints of the cost of the obtained biosensor, the ease of control, the simplicity of the device, the accuracy of temperature controllability, etc., a desired potential is applied to the resistor, or a desired current value is applied to generate heat to obtain the desired value. It is preferable to keep the temperature. In particular, when a printed resistor is used as the resistance element, it can be easily provided on the insulating substrate, and thus it is more preferable as the heating element.

As a method for providing the printed resistor, a resistor paste used for general electric and electronic circuits may be used by a method such as screen printing to print the resistor with a predetermined shape and a predetermined thickness to obtain a resistor. . As resistance paste,
Polymer-type thick film resistance paste using carbon, graphite, etc. as a conductive filler and thermosetting resin such as epoxy resin, phenol resin, etc. as binder, or cermet type obtained by sintering such as ruthenium oxide etc. and glass frit etc. Thick film paste or the like can be used. Both of these require a heating process after printing, but when a resistor is provided after forming the measurement electrode, a UV-curable or electron-beam curable carbon paste that does not require a heating process can be used. .

To provide the above-mentioned printed resistor, for example, the resistor is printed on the back surface of the insulating substrate provided with the electrode system, or the resistor is printed so that the resistor does not come into direct contact with the outside air. An appropriate method such as bonding another substrate to the insulating substrate can be used.

The measuring electrode is formed by forming a conductive layer containing an organic charge transfer complex crystal on a conductive substrate and immobilizing both the enzyme and the electron mediator. The organic charge transfer complex crystal may be grown in an insulating polymer film on a conductive substrate, but is preferably formed directly on the conductive substrate.

To grow an organic charge transfer complex crystal in an insulating polymer film to form a conductive layer, for example, an electron donor layer is provided on a conductive substrate, and polyvinyl butyral, polyester, There is a method in which a coating film of an insulating polymer such as polyamide or polyether amide is provided and this is brought into contact with a solution containing an organic electron acceptor. An enzyme or an enzyme and an electron mediator can be immobilized on this conductive layer to manufacture an enzyme electrode.

The conductive layer composed of the organic charge transfer complex crystal formed directly on the surface of the conductive substrate can be easily obtained by growing the crystal in the thickness direction by the following method or the like. It has excellent conductivity.

Examples of the conductive substrate include metals such as copper, silver, platinum and gold and carbon electrodes, a substrate provided with a conductive layer made of these conductive materials on the surface by means such as vapor deposition, or a conductive substrate of these. A substrate or the like prepared from a paste containing a powder of a conductive material can be used.

Here, the organic charge transfer complex (hereinafter, organic CT
The term "complex" is a compound formed from an organic electron acceptor and an electron donor by a charge transfer reaction between them. The organic electron acceptor used for forming the organic CT complex is not particularly limited, but a compound having a cyanomethylene functional group is preferable, and a compound having a dicyanomethylene functional group and a quinone or naphthoquinone skeleton is particularly preferable. Of these, 7,7 ', 8,8'-tetracyanoquinodimethane (TCNQ) has a strong CT complex-forming ability, and the resulting organic CT complex has a high electric conductivity, which is advantageous in response time and responsiveness. Is. Moreover, it is preferable because it is relatively easy to obtain industrially.

The electron donor used for forming the organic CT complex is not particularly limited as long as it can form a CT complex having conductivity with the organic electron acceptor used, and it may be organic or inorganic. Either of them is okay. Specifically, inorganic materials include copper, silver, cobalt, nickel,
Iron, manganese, etc., and as organic materials, tetracene compounds such as tetrathiafulvalene, tetraselenofulvalene, and their derivatives, or 2,2'-bispyridinium, N-methylphenazinium, etc., known electron donors. Can be used.

To grow organic CT complex crystals, known methods in liquid phase and vapor phase can be used. Organic in liquid phase
The following methods are available as methods for growing CT complex crystals. First, a part or all of a substrate such as a substrate provided with an electron donor layer on the surface thereof or a copper plate which also functions as an electron donor is brought into contact with a solution containing an organic electron acceptor. As a result, the organic electron acceptor in the solution causes a CT complexation reaction with the electron donor constituting the surface of the substrate, and the complex grows.

As the solvent used for preparing the organic electron acceptor-containing solution, a polar aprotic solvent such as acetonitrile, dioxane, N, N-dimethylformamide,
Dimethyl sulfoxide, hexamethylphosphoramide,
Methyl ethyl ketone and the like are preferable. The concentration of organic electron acceptor in this solution is usually 100 parts by weight of solvent.
0.01 parts by weight to saturation concentration, preferably 0.1 parts by weight to saturation concentration are suitable.

The formation of the organic CT complex is usually carried out at a temperature of 10 to 30 ° C., but depending on the combination of the organic electron acceptor used and the electron donor on the surface of the substrate, the CT complexation reaction proceeds rapidly, resulting in a dense structure. It may be difficult to obtain a uniform target layer. In such a case, the temperature of the solution, the substrate, the atmosphere may be lowered or the concentration of the solution may be lowered, if necessary. On the contrary, when the complexing reaction is slow and it takes a long time for the organic CT complex crystal to grow to the required thickness, heating can be performed as necessary.

The contact time of the organic electron acceptor-containing solution largely depends on the combination of the organic electron acceptor and the electron donor used and the thickness of the target conductive layer, but it is generally about 10 seconds to 1 hour. A vacuum deposition method can be generally used as the vapor phase growth method. First, a substrate provided with an electron donor layer or a copper plate that also functions as an electron donor is placed under reduced pressure (1 × 10 ⁻³ to 1 × 10 ⁻⁷ torr) to form a complex crystal. The part you want to grow should
Hold at ~ 300 ° C). Next, the electron acceptor is gradually heated and vaporized. As a result, the complex grows by the complexing reaction between the electron acceptor molecule reaching the surface of the substrate and the electron donor on the surface of the substrate. At this time, the thickness of the conductive layer can be easily controlled by the temperature of the substrate, the vaporization rate of the electron acceptor and the like.

The organic CT complex prepared by such a liquid phase method or a vapor phase method generally becomes fine needle crystals and grows in the direction perpendicular to the substrate surface. The thickness of the conductive layer composed of this organic CT complex is not particularly limited, but is usually in the range of 0.01 to 50 μm, preferably 0.1 to 10 μm.
Is.

As described above, the conductive layer is composed of fine needle-like crystals grown in the thickness direction thereof, so that the surface of the conductive layer has a structure having fine irregularities. Therefore, when immobilizing an enzyme or an electron mediator, which will be described later, the enzyme or the electron mediator can be captured in the fine recesses, and the immobilization thereof becomes easy. In addition, since it is a fine needle crystal, the actual surface area of the electrode part can be made large, and as a result, the amount of enzyme and electron mediator immobilized is increased, and the current density obtained as the output of the enzyme electrode is increased. It becomes possible.

If the thickness of this conductive layer is too thin within the above range, a sufficient surface area cannot be obtained, and as a result, the output current value becomes small. On the other hand, when the thickness is more than the above range, the resistance of the conductive layer itself becomes large. Therefore, when used as an enzyme electrode, a voltage drop occurs on the electrode surface when a voltage is applied. Also this organic
Since the CT complex itself does not have high mechanical strength, if it is too thick, structural defects are likely to occur.

The conductive layer is washed and dried, if necessary, and then an enzyme consisting of a water-insoluble polymer and an enzyme and an electron mediator are immobilized on the electrode composed of the conductive substrate and the conductive layer so as to be in contact with the conductive layer. Alternatively, when reusability is not required, such as disposable, the enzyme and electron mediator may be immobilized without using this water-insoluble polymer.

The enzyme can be appropriately selected according to the target substance and the target chemical reaction in consideration of the substrate specificity and reaction specificity of the enzyme. The enzyme that can be used is not particularly limited, and examples thereof include glucose oxidase, alcohol dehydrogenase, peroxidase, catalase, lactate dehydrogenase, galactose oxidase, penicillinase and the like. A combination of oxidoreductase and coenzyme is also possible.

The electron mediator to be used may be one that can efficiently carry out the electron transfer accompanying the enzymatic reaction, that is, that can smoothly carry out the electron transfer from the enzyme to the organic CT complex. For example, in the case of a reaction in which a substrate is oxidized by an oxidase, an electron is easily accepted from the reduced enzyme, the electron mediator itself is reduced, and the electron is donated to the electrode by an electrode reaction on the surface of the conductive layer. , Which has the property of returning to the oxidized form. Such electron mediators include ferrocene, 1,1′-dimethylferrocene, ferrocenecarboxylic acid, ferrocene derivatives such as ferrocenecarboxaldehyde, quinones such as hydroquinone, chloranil, bromanil, ferricyan ion, octacyanotungstate ion, A metal complex ion such as octacyanomolybdate ion is suitable.

As a method for immobilizing an enzyme and an electron mediator on a conductive layer made of an organic CT complex by using a water-insoluble polymer, the following method is possible. (1) A solution containing an enzyme and an electron mediator on the conductive layer is coated and dried, and then a water-insoluble polymer solution is coated and dried to form a conductive layer on the conductive substrate,
An intermediate layer composed of an enzyme and an electron mediator and three layers composed of a water-insoluble polymer coating layer are provided. (2) by coating a water-insoluble polymer solution containing an enzyme and an electron mediator on the conductive layer, and drying,
A conductive layer and a water-insoluble polymer coating layer containing an enzyme and an electron mediator are provided on the conductive substrate. (3) applying a solution containing an enzyme on the conductive layer, and drying,
Then, an electron mediator, or a water-insoluble polymer solution containing the electron mediator and the enzyme is applied and dried to form a conductive layer on the conductive substrate, an intermediate layer composed of the enzyme, and a water containing the electron mediator or the electron mediator and the enzyme. Insoluble polymer coating layer 3
Provide layers. (4) A solution containing an electron mediator on the conductive layer is coated and dried, and then a water-insoluble polymer solution containing an enzyme or an enzyme and an electron mediator is coated and dried to form a conductive layer on the conductive substrate. , An intermediate layer composed of an electron mediator and a water-insoluble polymer coating layer containing an enzyme or an enzyme and an electron mediator are provided.

For immobilizing enzymes and electron mediators, the above methods are simple and suitable, but are not limited thereto, and known covalent bonding methods, ionic bonding methods, adsorption methods, entrapment methods, cross-linking methods and the like can be used. It is also possible to use.

As the water-insoluble polymer, a film can be easily and uniformly formed, and an enzyme and an electron mediator are uniformly dispersed and fixed, and when used as an enzyme electrode, the enzyme is dissolved and swelled in a sample solution to form an enzyme. Any material can be used without limitation as long as it does not cause a decrease in output due to elution of the electron mediator. Furthermore, if pinholes are formed in the conductive layer, when used as an enzyme electrode, the substrate or electron donor may be eluted into the sample solution.
The water-insoluble polymer layer covers the pinhole portion to prevent elution.

Examples of such water-insoluble polymers include thermoplastic polymers such as polyvinyl butyral, polyesters, polyamides and polyesteramides.
One kind or two or more kinds can be used. Depending on the method of immobilizing the enzyme and the electron mediator, and in consideration of the diffusibility of the substrate, etc., the polymer can be appropriately selected. For example, polymer vinyl butyral has water-insolubility, hydrophilicity, and water absorption, and It is suitable because it has very micropores.

To coat the conductive layer with a water-insoluble polymer containing an enzyme and an electron mediator, the enzyme and the electron mediator are dissolved or uniformly dispersed in a solution prepared by dissolving the water-insoluble polymer in a suitable organic solvent. Then, this can be carried out by directly coating and drying the conductive layer. When the enzyme and the electron mediator are dispersed and used, the enzyme and the electron mediator are improved in dispersion and solubility by appropriately adding water, which is a good solvent for the enzyme and the electron mediator, within a range in which the water-insoluble polymer does not precipitate. Can be fixed efficiently. The obtained enzyme electrode is
After washing with pure water or a buffer solution to remove the enzyme and electron mediator which are not completely immobilized, the product can be used. Further, it is advantageous to grow the organic CT complex by a means such as further immersing the enzyme electrode in an electron acceptor solution, because the conductivity of the entire membrane is increased and a large response current can be obtained.

As described above, when the conductive layer is coated with the water-insoluble polymer containing the enzyme and the electron mediator, a smooth transition from the enzyme to the organic CT complex, from the enzyme to the electron mediator, or from the electron mediator to the organic CT complex is performed. In order to secure electron mobility and improve responsiveness, the fixing film should be thin. For example, it is 10Å-10 μm, preferably 100Å-1 μm. Further, it is not necessarily required that the film is uniform, and the enzyme and the organic CT complex, the enzyme and the electron mediator, or the electron mediator and the organic CT complex may be brought into direct contact with each other.

When the intermediate layer consisting of the enzyme and the electron mediator is coated with the water-insoluble polymer, 0.01 to 10 μm is preferable in consideration of contact between the enzyme and the substrate and coating of the remaining pinholes. It is desirable that the thickness is about 0.1 to 5 μm. The enzyme electrode thus obtained is
It has a structure in which an enzyme and an electron mediator are fixed so that they are in contact with each other on a conductive layer provided on a conductive substrate.Since it is structurally simpler than conventional hydrogen peroxide electrodes, oxygen electrodes, etc., it can be miniaturized. is there. In addition, the conductive layer made of the organic CT complex crystal not only facilitates electron transfer with the enzyme,
The electric conductivity of the crystal layer is remarkably higher than that of ferrocene or the like which has been conventionally used as an electron mediator. It is considered that this is because the organic CT complex forms a needle-shaped crystal in which it has developed, so that even with the same content, the number of conductive paths in the film increases, which effectively contributes to electron transfer. In addition, since the electron mediator is immobilized on the surface of the conductive layer, smooth electron mobility is ensured even in the area where electron transfer does not easily occur even though the organic CT complex is in contact with the enzyme. It is possible to improve the responsiveness. Moreover, when organic CT complex crystals are grown directly on a conductive substrate without using a polymer, a fine uneven surface of needle-like crystals is obtained, so the amount of enzyme or electron mediator that directly contacts the conductive layer is increased. Therefore, the responsiveness of the enzyme electrode can be further enhanced.

The counter electrode has a low resistance of the electrode itself so that when the constant potential is applied to the measuring electrode or the compensating electrode, the current in these electrodes flows without hindrance. It is not used, and the reaction product at the counter electrode does not interfere with the reaction at the measurement electrode, and the reaction product itself does not react. From such a viewpoint, a metal such as platinum, gold, silver, and copper, a carbon electrode, or a conductive layer made of these conductive materials may be provided on the surface of the substrate by a method such as vapor deposition or sputtering, or the conductivity of these may be provided. What was obtained by making from the paste containing the powder of the conductive material can be used as a counter electrode. Further, for example, when silver is used, it is preferable that AgCl is deposited on the surface by means of anodic polarization or the like and electrochemically stabilized in a solution.

The counter electrode and the measuring electrode can be provided in various shapes, but the method of providing them on the same plane is preferable from the viewpoints of manufacturing cost, process simplicity, sample volume required for measurement, and the like. When they are provided on the same support, a metal layer of copper, silver, gold, mercury or the like is formed on the insulating substrate by vapor deposition, sputtering, etc. to form the conductive substrate and the counter electrode in the measurement electrode. Further, the organic CT complex formed directly on the conductive substrate at the measuring electrode may be similarly grown on the counter electrode and used as the counter electrode. In this case, in the structure in which the measurement electrode and the counter electrode are arranged on the same support, the counter electrode can be simultaneously prepared in the step of growing the organic CT complex of the measurement electrode, which is an advantage that the process is simple.

The resistance of the compensating electrode itself is about the same as the resistance of the measuring electrode, and it does not react specifically with proteins, electrolytes, biological components, etc. in the sample solution, and the electrochemical other than the enzymatic reaction at the measuring electrode The shape is preferably the same or at least the same area for the purpose of compensating for the current value, although it is not particularly limited as long as it is affected by the secondary side reaction or adsorption to the same extent. . When the electrode system is provided on the same support, the compensation electrode can be prepared at the same time as the preparation of the measurement electrode except that the enzyme is immobilized. For example, directly on the conductive substrate at the measuring electrode, organic CT
The compensating electrode can be simultaneously prepared in the step of forming the complex and the step of applying the water-insoluble polymer, which simplifies the process.

The measurement using the biosensor of the present invention can be performed without using the reference electrode. In such a case, the area of the counter electrode is more than twice the area of the measurement electrode,
It is preferably 10 times or more. This is because the potential difference applied at the time of measurement is mainly applied to the measurement electrode, so that the measurement can be performed with high accuracy.

In the present invention, a filtrate layer and / or a liquid retaining layer may be further provided on the electrode system obtained as described above. The filtrate layer is provided for the purpose of reducing the influence of relatively large interfering substances in the sample. The liquid retaining layer is provided to hold a small amount of sample and to uniformly diffuse the sample. These may be provided individually for each purpose, but it is preferable to provide both layers at the same time. Further, the filtrate layer and / or the liquid retaining layer may be installed so as to cover at least the measurement electrode, and may cover only the measurement electrode or the measurement electrode and the counter electrode. When the compensation pole is provided, the compensation pole may be covered. When both the filtrate layer and the liquid-retaining layer are provided, it is preferable to provide the liquid-retaining layer and then the filtrate layer in this order on the electrode.

As the liquid-retaining layer, one having absorptivity for a small amount of supplied sample and capable of uniformly diffusing the sample on the electrode system is used. From such a viewpoint, a water-absorbing polymer is preferable in the case of an aqueous-system biological sample, and carboxymethylcellulose, gelatin, methylcellulose, starch-based or acrylate-based, vinyl alcohol-based, vinylpyrrolidone-based or maleic anhydride is used. Polymers of the system are preferred. Since these polymer substances can be easily made into an aqueous solution, a thin film having a required thickness can be formed by applying an aqueous solution having an appropriate concentration onto the electrode and drying. Also, using a hydrophilic non-woven fabric such as rayon,
It is also possible to supply the supply sample to the electrode by utilizing the capillary phenomenon. When a hydrophilic non-woven fabric is used, it is provided on the electrode by a method such as sticking it with a hydrophilic adhesive or mechanically setting it. The thickness of the liquid retaining layer is about 0.1-10 μm.

The filtrate layer has a function of suppressing diffusion of relatively large interfering substances such as cells, erythrocytes, and proteins contained in the sample to be supplied onto the electrode surface, and allowing the substrate to be measured to permeate. Is. As a material having such a function, a porous material such as polycarbonate, cellulose acetate, nylon non-woven fabric, rayon, or ceramics is suitable, and its pore diameter is usually 0.001 μm.
.About.10 .mu.m, preferably 0.01 .mu.m to 1 .mu.m is used. The formation of the filtrate layer is performed by applying a solution of a polymer substance on the electrode or the liquid-retaining layer and drying it, or by setting it by a method such as mechanically adhering a thin film of a porous body. be able to.

In a biosensor provided with a filtrate layer and / or a liquid retaining layer, the sample can be dropped directly onto the filtrate layer or the liquid retaining layer for measurement. When both layers are provided, interfering substances are removed in the filtrate layer and the sample that has penetrated into the liquid retention layer diffuses uniformly on the electrode, so even a small amount of sample can be used as it is for accurate measurement. .

In the filtrate layer and the liquid-retaining layer, if an anticoagulant is carried in at least one of these layers, the whole blood as collected as a measurement sample can be used for quick and stable measurement. You can Blood has a high viscosity, and it took one minute or more to drop the electrode, pass through the filtrate layer or the liquid-retaining layer, reach the enzyme immobilization layer, and obtain an output current corresponding to the substrate concentration. It is considered that this is because blood begins to coagulate from the moment it comes into contact with the atmosphere and gradually thickens. When an anticoagulant is used, it reaches the enzyme immobilization layer in a short time of, for example, about 15 seconds and a constant output current value is obtained. As an anticoagulant to be used, sodium fluoride is preferable because it is stable and easy to handle, but in addition, heparin, sodium citrate, and ethylenediaminetetraacetic acid can also be used to shorten the transit time of the sample solution and The effect that a response current is obtained is exhibited. As a method of supporting the anticoagulant, there is a method of using a solution in which the anticoagulant is dissolved or dispersed at a predetermined concentration in the coating solution when forming the filtrate layer or the liquid retaining layer.

When a potential is applied to the biosensor of the present invention to measure a response current due to an enzymatic reaction, it is preferable to apply a pulse potential. When a steady-state current is applied, the surface state of the electrode changes due to an electrochemical reaction other than the purpose, so that a measurement error is likely to occur. It is preferable to apply a pulse potential because such deterioration of the electrode can be minimized and the stabilization time is short. Further, in the measurement using a biosensor provided with a compensation electrode, a predetermined potential is continuously or simultaneously applied between the measurement electrode and the counter electrode, and between the compensation electrode and the counter electrode, whereby the current flowing between both electrodes is increased. The value can be measured, and the difference in current value between the two can be quantified as the response current due to the enzymatic reaction.

By using the biosensor of the present invention, it is possible to measure sugars such as glucose, lactic acid, alcohol and other trace biological substances in blood and urine, and sugars and alcohols in food processing processes. In the biosensor of the present invention, a sample such as blood can be used as it is without diluting or agitating the sample, and even if the sample is a small amount, it can be easily analyzed with high accuracy. Stable measurement is possible without being affected by temperature and sample solution temperature.

[0052]

【Example】

[0053]

Example 1 A double-sided copper-clad glass epoxy substrate (R-1701 manufactured by Matsushita Electric Works, Ltd.) was etched to form a measurement pole portion (diameter 1.5 mm) and a counter pole portion having the shape shown in FIG. 1, and the back side of FIG. A terminal portion was formed and the entire surface was plated with electrolytic silver to obtain an electrode. Next, on the front side, an epoxy resin paint was screen-printed and molded, leaving the electrode portion. On the back side, a polymer type thick film resistor paste was printed in the shape shown in FIG. 2 and cured to obtain a printed resistor (area resistance of about 1 KΩ / □) having a thickness of about 20 μm. Of the electrode part on the front side, the counter electrode part
Anodic polarization was performed for 2 minutes at a current density of 0.4 mA / cm ^-2 in 0.1 M hydrochloric acid to deposit silver chloride on the surface.

1.0 g of 7,7 ', 8,8'-tetracyanoquinodimethane (reagent, manufactured by Kishida Chemical Co., Ltd., abbreviated as TCNQ hereinafter) was added to 10 mg of acetonitrile (for reagents and spectra) to give a saturated solution of TCNQ. Prepared. Weigh 2 μl of this TCNQ saturated solution,
At room temperature, it was dropped on the measurement electrode part of the above electrode and naturally dried. This operation was repeated 5 times in total to form an organic CT complex thin film having dark purple fine needle-shaped crystals on the entire surface of the measurement electrode.

Glucose oxidase (Aspergillus n
iger-derived, Sigma, Type VII) 40mg 1ml 100m
After dissolving in M phosphate buffer (pH 7.0), 4 μl was measured and applied to the above-mentioned measurement electrode part and air-dried.

Polyvinyl butyral resin [trade name: S-REC B, BX-L, manufactured by Sekisui Chemical Co., Ltd.] 1.0 g and 1,1'-dimethylferrocene [reagent, Tokyo Kasei Co., Ltd.]
Manufacture] 1.0 g was dissolved in 2.5 g of ethyl carbitol to prepare a composition containing a reduced mediator. The composition was printed on the measurement electrode portion using a screen printer and dried at 80 ° C. for 10 minutes to form a layer containing a reduced mediator having a thickness of about 3 μm.

Next, 1 μl of the above-mentioned TCNQ solution was measured,
After coating on the above-mentioned measurement electrode and air-drying, it was washed with pure water.
5 μl of an aqueous solution containing 1% by weight of carboxymethylcellulose and 2% by weight of sodium fluoride was measured and applied on the measurement electrode and the counter electrode and dried, and then a methylene chloride solution of polycarbonate was applied and dried to prepare a sensor. did.

A potential was applied to the terminal portion of the resistor on the back surface, and the surface temperature of the electrode surface was maintained at about 40 ° C. while adjusting the potential.
In this state, this sensor was placed in each atmosphere temperature shown in FIG. 3, and 5 μl of a whole blood sample having a glucose concentration of 10 mM was dropped on the electrode system to cover the entire surface of the measurement electrode and the counter electrode. After leaving it as it is for 1 minute, 0.25V is applied to the counter electrode.
Was applied to the measurement electrode, and the current value 5 seconds after the application of the potential was measured. The results obtained in the same manner at each ambient temperature are shown in FIG. As described above, it was found that the biosensor of the present invention is less affected by the change in ambient temperature and can be measured with high accuracy.

[0059]

Comparative Example 1 FIG. 3 shows the results of measurement performed in the same manner as in Example 1 except that no electric potential was applied to the resistor on the back surface of the electrode. In this case, the response current changes greatly with changes in the ambient temperature field.

[0060]

Example 2 A single-sided copper-clad flexible polyimide substrate was etched to prepare an electrode pattern having a measuring electrode, a counter electrode, and a compensating electrode having the shape shown in FIG. 4, and the counter electrode was prepared in the same manner as in Example 1. After the conductive organic charge transfer complex thin film was formed on the measurement electrode part and the compensation electrode part in the same manner as in Example 1, the enzyme solution was applied only to the measurement electrode part in the same manner.

5.0 g of TCNQ was suspended in 300 ml of acetone at room temperature, an acetone solution of TCNQ and an equivalent amount of dimethylferrocene was gradually added dropwise, and the mixture was stirred for another 2 hours and then filtered,
It was dried to obtain a dimethylferrocene-TCNQ complex.

Polyvinyl butyral resin (1.0 g), dimethylferrocene-TCNQ complex (0.7 g) and TCNQ (0.3 g) were weighed and added with 2.5 g of ethyl carbitol. It was made. This is printed in the same manner as in Example 1 by the screen printing method on the measurement pole portion and the compensation pole, and dried,
A layer containing oxidized media having a thickness of about 4 μm was prepared on both electrodes.

Further, an aqueous solution containing 3% by weight of polyvinyl alcohol and 2% by weight of heparin was applied as a liquid-retaining layer to the entire surface of the electrode system including the measuring electrode, and after drying, an acetone solution of cellulose acetate was used as a filtrate layer. Was applied and dried to prepare a sensor.

Separately from this, a terminal portion having a shape shown in FIG. 5 was formed on a glass plate with silver paste, and a polymer type thick film resistor paste was printed in a shape shown in FIG. 5 (resistor film thickness: about 20 μm). . A sensor was formed by laminating the electrode system on the resistor so that the electrode system was mounted on the resistor.

Using this sensor, the temperature of the electrode surface was set to about 40 ° C. while adjusting the applied potential in the same manner as in Example 1, and the sensor was placed within the ambient temperature shown in FIG. 5 μl of a whole blood sample having a glucose concentration of 20 mM was dropped on the electrode system, and the entire surface of the electrode system was covered with the sample solution through the liquid retaining layer. As it is
After standing for 30 seconds, a pulse potential of 0.25 V was applied to the measurement electrode against the counter electrode, and the current value 5 seconds after the application was measured. Subsequently, a pulse potential of 0.25 V was applied to the compensating electrode with respect to the counter electrode, and the current value 5 seconds after the application was similarly measured. The difference between the current values obtained at the measurement pole and the compensation pole was taken as the response current. As shown in FIG. 6, when the biosensor of the present invention was used, a substantially constant response current was obtained without being affected by the ambient temperature.

[0066]

COMPARATIVE EXAMPLE 2 The result of measurement in the same manner as in Example 2 without applying a potential to the resistor on the glass plate in Example 2 is shown in FIG. In this case, the response current changes greatly with changes in the ambient temperature.

[0067]

INDUSTRIAL APPLICABILITY According to the present invention, a biosensor that uses an enzyme electrode using an organic charge transfer complex as an electrode material and can easily measure even a small amount of sample with high accuracy over a long period of time is further provided. By providing the element, it is possible to provide a biosensor that can always perform stable and highly accurate measurement without being affected by the outside air temperature and the sample liquid temperature. Further, if a filtrate layer or a liquid retention layer is provided on the electrode, it has good responsiveness even in the case of a high-concentration substrate, and it becomes possible to measure an extremely small amount of sample. Furthermore, if at least one layer of the filtrate layer and the liquid-retaining layer is loaded with an anticoagulant, more rapid measurement can be performed.

[Brief description of drawings]

FIG. 1 is a diagram showing an electrode surface of a biosensor of the present invention.

FIG. 2 is a view showing the back surface of the electrode surface of the biosensor of the present invention.

FIG. 3 is a diagram showing the relationship between the ambient temperature and the response current measured in Example 1 and Comparative Example 1.

FIG. 4 is a diagram showing another example of the electrode surface of the biosensor of the present invention.

FIG. 5 is a diagram showing another example of the back surface of the electrode surface of the biosensor of the present invention.

FIG. 6 is a diagram showing the relationship between the ambient temperature and the response current measured in Example 2 and Comparative Example 2.

[Explanation of symbols]

 1: Substrate 2: Counter electrode 3: Measurement electrode 4: Compensation electrode 5: Epoxy resin 6: Lead part 7: Counter electrode terminal 8: Measurement electrode terminal 9: Compensation electrode terminal 10: Printing resistor 11: Resistor terminal

Claims (5)

[Claims] 1. An electrode system comprising a measuring electrode and a counter electrode is provided on one surface of an insulating substrate, and the measuring electrode has an enzyme and an electron mediator immobilized on an electrode containing an organic charge transfer complex crystal as a conductive layer. In a certain biosensor, a heating element is provided on the other surface side of the insulating substrate or in the insulating substrate. 2. The biosensor according to claim 1, wherein a filtrate layer and / or a liquid retaining layer is provided on the measurement electrode or on the electrode system including the measurement electrode. 3. The biosensor according to claim 2, wherein an anticoagulant is carried in the filtrate layer and / or the liquid retaining layer. 4. The biosensor according to claim 1, further comprising a compensation electrode having no enzyme immobilized thereon. 5. The biosensor according to claim 1, wherein the enzyme is a redox enzyme.