JPH05203608A - Biosensor - Google Patents

Abstract

PURPOSE:To measure sample of a trace quantity without diluting or agitating it by providing an electrode system comprising a measuring pole at which enzyme having a conductive layer including organic charge-transfer complex crystals is immobilized, and a counter pole provided close to it. CONSTITUTION:A biosensor comprises at least a measuring pole 3 and a counter pole 2 provided close to it, and the measuring pole 3 is an electrode including organic charge-transfer complex, or organic charge-transfer complex and electron mediator, having immobilized enzyme. A compensation pole 4 at which enzyme is not immobilized may be provided additionally for achieving high precision measuring without having effects of adsorption of protein, etc., in sample or other auxiliary reaction. In case an electrode system is provided in the same support body, the compensation pole 4 can be manufactured simultaneously and similarly to manufacture of the measuring pole 3 except for immobilizing enzyme. For example, the compensation pole 4 can be manufactured simultaneously in a process of forming organic CT complex on conductive substrate and a process of applying water-unsoluble high polymer for the measuring pole 3 directly, thereby the processes can be simplified.

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 for measuring the concentration of a trace amount of a biological substrate contained in body fluid components such as blood and urine, and a method for using the enzyme sensor.

[0002]

2. Description of the Related Art Analytical methods utilizing the excellent substrate specificity of enzymes have attracted attention in the fields of clinical analytical chemistry, food manufacturing, environmental chemistry and the like. In particular, in the field of clinical analytical chemistry, enzyme electrodes capable of selectively detecting glucose, urea, uric acid and the like have been conventionally known. These enzyme electrodes are generally composed of an electrode and an enzyme-immobilized membrane, and the substance change due to the enzyme reaction is read by the electrode as a change amount of an electric signal to measure the concentration of a substrate on which the enzyme specifically acts. It is a thing. For example, in a glucose sensor,
The biological substrate concentration is measured by monitoring the electrode active substances such as hydrogen peroxide and oxygen produced or consumed by the enzymatic reaction with the electrode as shown in the following formula.

[0003]

[Chemical 1]

However, the enzyme electrode based on such a principle has the following problems.

As is clear from the above equation, the reaction of the substrate requires stoichiometric oxygen, but in actual measurement, for example, a glucose sensor measures the blood glucose concentration in a diabetic patient. In this case, the amount of dissolved oxygen in body fluid is insufficient. Therefore, dilute the sample blood,
Measures such as supplementing oxygen in some way are taken.

In the case of electrically monitoring hydrogen peroxide, when a substance that is oxidized at the same potential as hydrogen peroxide, such as a reducing substance such as ascorbic acid, is present in the sample solution, these interfering substances are present in the measurement current. Oxidation current is added, resulting in a measurement error. Therefore, in order to eliminate these errors, the electrode on which the enzyme is not fixed is corrected as a compensation electrode, and oxygen, hydrogen peroxide molecules and the measurement substrate are permeated, but an electrode active buffer substance such as ascorbic acid is not permeated. It was necessary to attach a permselective membrane such as.

As described above, the sensor based on the principle of measuring the concentration of the substance produced or consumed by the enzymatic reaction has problems such as the influence of dissolved oxygen and the influence of interfering substances. Further, there is a limit to miniaturization because it requires a form in which an enzyme-immobilized membrane is attached to oxygen and hydrogen peroxide electrodes.

On the other hand, in order to solve these problems, an enzyme electrode using a conductive polymer and an enzyme electrode using an electron mediator have been proposed. In the former case, during electropolymerization of conductive polymers such as polypyrrole and polyaniline,
The enzyme is allowed to coexist in the monomer solution, and the enzyme is trapped in the polymerized film at the time of polymerization, or the enzyme-immobilized film is provided on the pre-polymerized conductive polymer film by a known method to form a conductive enzyme-immobilized film. I will get it. Further, an enzyme electrode using an electron mediator, ferrocene, benzoquinone, ferricyanide ion, N-methylphenazinium and other electron mediators are contained in the carbon paste, and the enzyme is immobilized on the surface of the carbon paste electrode. It is coated with a suitable polymer film. However, an enzyme electrode that directly detects electron transfer associated with an enzymatic reaction using a conductive polymer has the advantage that it is not affected by dissolved oxygen, but it has low responsiveness and long response time. There is. Furthermore, when the method of trapping 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. Further, the conductivity is low even in the enzyme electrode using the conventional electron mediator, and the response is insufficient, and in addition, the response time is insufficient, and since the electron mediator has a form dispersed in the carbon paste, the elution of the electron mediator, There is a problem that the responsiveness decreases with time due to desorption.

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 measure with high accuracy, but also operation such as dilution and agitation of the sample solution is 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 often restrictions on the amount of sample that can be used, and measurement with a small amount of sample is desired.
No conventional enzyme electrode has been able to perform simple and quick and accurate measurement for a long time even with a small amount of sample.

Therefore, as a solution to these drawbacks of the conventional enzyme electrodes, the present inventors have 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. 2-24484). This enzyme electrode adopts a method of directly detecting the electron transfer accompanying the enzyme reaction, and thus has the advantage that it is not affected by dissolved oxygen and is less affected by interfering substances, and that it has excellent stability over time,
It has an advantage that a highly accurate response can be given over a long period of time. 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). However, there is no need for operations such as dilution and stirring of the sample liquid, and further improvement is desired for enzyme electrodes suitable for measurement with a small amount of sample.

[0011]

DISCLOSURE OF THE INVENTION The present invention provides a biosensor capable of measuring a small amount of sample as it is without diluting or stirring in a measuring system including an enzyme electrode using an organic charge transfer complex. With the goal.

[0012]

The inventors of the present invention directly measure a sample by miniaturizing an enzyme-immobilized electrode prepared by using an organic charge transfer complex as an electrode material and disposing a counter electrode in the vicinity thereof. As a result, they have found that a small amount of sample may be used, and have completed the present invention. Furthermore, the present inventors formed a conductive layer composed of an organic charge transfer complex on a conductive substrate, and contacted with this conductive layer to immobilize or coat the enzyme and the electron mediator with or without a water-insoluble polymer. It was also found that it is more preferable to use the above electrode as the measurement electrode.

The gist of the present invention is a biosensor having an electrode system having a measuring electrode on which an enzyme having a conductive layer containing an organic charge transfer complex crystal is immobilized and a counter electrode provided in the vicinity of the measuring electrode. Further, the present invention provides the biosensor,
It also relates to a biosensor provided with a compensating electrode in which an enzyme is not immobilized. As a preferred embodiment of the biosensor of the present invention,
As a measuring electrode, an electrode having a conductive substrate and a conductive layer made of an organic charge transfer complex crystal provided on the surface of the substrate and having an enzyme and an electron mediator immobilized with or without a water-insoluble polymer was used. It is preferable to use, and in this case, the compensation electrode has a conductive substrate and a conductive layer made of an organic charge transfer complex crystal provided on the surface of the substrate,
It is preferable to use an electrode on which an electron mediator is immobilized, with or without a water-insoluble polymer. The enzyme electrode of the present invention is particularly suitable when an oxidoreductase is used. Furthermore, the present invention also relates to a method for analyzing a specific substance by applying a pulse potential to the above biosensor and measuring a response current due to an enzymatic reaction.

[0014]

The biosensor of the present invention comprises at least a measurement electrode and a counter electrode provided in the vicinity thereof, and the measurement electrode is an electrode containing an organic charge transfer complex, or an organic charge transfer complex and an electron mediator, on which an enzyme is immobilized. is there. Further, in the biosensor of the present invention, it is possible to provide a compensating electrode in which an enzyme is not immobilized and to perform highly accurate measurement without being affected by adsorption of proteins and the like in samples and side reactions.

This enzyme-immobilized electrode is one in which a conductive layer containing an organic charge transfer complex crystal is formed on a conductive substrate to immobilize an enzyme or both an enzyme and an 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 polyethylene vinyl butyral or polyester is provided thereon. There is a method in which a coating film of an insulating polymer such as polyamide, polyesteramide or the like is provided and is brought into contact with a solution containing an organic electron acceptor. An enzyme or an enzyme and an electron mediator are immobilized on this conductive layer to obtain 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 good 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 the like. A substrate or the like prepared from a paste containing powder can be used.

Here, an organic charge transfer complex (hereinafter referred to as 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 preferable. Is. Among them, 7,7 ', 8,8'-
Tetracyanoquinodimethane (TCNQ) has a strong CT complex-forming ability, and the obtained organic CT complex has a high electric conductivity, which is advantageous in response time and responsiveness. Further, 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. You can use either of these. Specifically, inorganic materials include copper, silver, cobalt, nickel,
Iron, manganese and the like, and as organic materials, tetracene and its derivatives such as tetrathiafulvalene and tetraselenofulvalene, or a known electron donor such as 2,2′-bispyridinium and N-methylphenazinium. Can be used.

To grow organic CT complex crystals, known methods in liquid phase and gas phase can be used. Organic in liquid phase
For example, the following method is available as a method for growing a CT complex crystal. First, part or all of the substrate such as the one provided with an electron donor layer on the surface of the substrate or the 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 to grow the organic CT complex crystal 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. Generally, a vacuum vapor deposition method can be 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 is at an appropriate temperature (100
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 substrate temperature, the vaporization rate of the electron acceptor and the like.

The organic CT complex prepared by the liquid phase method or the 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 the 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-like 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 the 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 value 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.

Direct organic deposition on the conductive substrate as described above.
A CT complex crystal is grown to form a conductive layer. This conductive layer is washed and dried as needed. An enzyme composed of a water-insoluble polymer is immobilized or coated on the electrode composed of the conductive substrate and the conductive layer so that the enzyme and the electron mediator are in contact with the conductive layer. Alternatively, the enzyme and the electron mediator may be immobilized without using the water-insoluble polymer when the reusability is not required such as disposable.

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 electron transfer accompanying the enzymatic reaction, that is, that can smoothly carry out 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, hydroquinone, chloranil, quinones such as bromanil, ferricyanide, octacyanotungstate ion, A metal complex ion such as an octacyanomolybdate ion is suitable. As a method of immobilizing or coating an enzyme or 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 is dropped on the conductive layer, dried, and then a water-insoluble polymer solution is applied and dried to form a conductive layer, an enzyme and an electron mediator on the conductive substrate. 3 intermediate layers consisting of 1 and a water-insoluble polymer coating layer are provided. (2) By applying a water-insoluble polymer solution containing an enzyme and an electron mediator on the conductive layer, and drying,
Two layers of a conductive layer and a water-insoluble polymer coating layer containing an enzyme and an electron mediator are provided on a conductive substrate. (3) A solution containing an enzyme is dropped on the conductive layer and dried,
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 is dropped on the conductive layer, dried, and then a water-insoluble polymer solution containing an enzyme or an enzyme and an electron mediator is applied and dried to form a conductive layer on a 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.

As described above, for immobilizing or coating the enzyme and the electron mediator, a polymer prepared by dissolving or uniformly dispersing the enzyme and the electron mediator in a water-insoluble polymer solution is directly applied to the conductive layer and dried to form the polymer. The method of coating the layer obtained by including in the solution or dropping the solution containing the enzyme or the electron mediator on the conductive layer and drying it with a water-insoluble polymer coating is simple and suitable, but is not limited to these. It is also possible to use a known covalent bond method, ionic bond method, adsorption method, entrapment method, crosslinking method or the like.

As the water-insoluble polymer, a uniform film can be easily formed, and an enzyme and an electron mediator are uniformly dispersed and fixed, and when it is used as an enzyme electrode, it dissolves and swells 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, polyester, polyamide and polyester amide.
One kind or two or more kinds can be used. Depending on the method of immobilizing the enzyme and the electron mediator, and the polymer can be appropriately selected in consideration of the diffusibility of the substrate, for example, polymer vinyl butyral is water-insoluble but has hydrophilicity and water absorption, and It is suitable because it has very micropores.

To coat the conductive layer with the water-insoluble polymer containing the enzyme and / or electron mediator, the enzyme and / or electron mediator is dissolved in a solution prepared by dissolving the water-insoluble polymer in a suitable organic solvent. Alternatively, it can be carried out by uniformly dispersing and directly coating and drying this on the conductive layer. When the enzyme and the electron mediator are used in a dispersed state, the dispersion of the enzyme and the electron mediator can be improved 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 can be used after being washed with pure water or a buffer solution to remove the enzyme and the electron mediator which are not completely immobilized. In addition, if the organic CT complex is grown by means such as further immersing the enzyme electrode in an electron acceptor solution, the conductivity of the entire film is increased,
It is advantageous in that a large response current can be obtained.

When the enzyme and / or the electron mediator is immobilized using a water-insoluble polymer, the enzyme is removed from the organic CT.
In order to ensure smooth electron transfer from the complex or enzyme to the electron mediator, or from the electron mediator to the organic CT complex to improve the 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 water-insoluble polymer is used to coat the enzyme and / or electron mediator-containing conductive layer or the intermediate layer containing the enzyme and / or electron mediator, contact between the enzyme and the substrate and the remaining pin Taking into account the hole coverage, 0.01-10 μm, preferably 0.1-
It is desirable to set the thickness to about 5 μm. The enzyme electrode thus obtained has a structure in which an enzyme and an electron mediator are fixed on a conductive layer provided on a conductive substrate so that the enzyme and the electron mediator are in contact with each other, and is structurally simpler than conventional hydrogen peroxide electrodes, oxygen electrodes, etc. Therefore, the size can be reduced. In addition, the conductive layer made of an organic CT complex crystal not only facilitates electron transfer with an enzyme, but also has an electric conductivity higher than that of ferrocene, which has been conventionally used as an electron mediator. Is significantly larger. It is considered that this is because the needle-like crystals in which these organic CT complexes are developed constitute the needle-like crystals, 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 transfer is ensured even in the part where electron transfer is structurally difficult even though the organic CT complex is in contact with the enzyme. It is possible to improve the responsiveness. Further, when the organic CT complex crystal is directly grown on the conductive substrate without using a polymer, a fine uneven surface of needle-like crystal is obtained, so that 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 biosensor of the present invention comprises a measuring electrode and a counter electrode, or a compensating electrode combined into a single sensor, which can be miniaturized. Further, the measuring electrode and the counter electrode, or even a compensating electrode can be provided. Since it can be measured by covering it with a sample, it is easy to measure even a small amount of sample by directly dropping it on the sensor.

The counter electrode has a low resistance of the electrode itself so that when a 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 materials 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.

The counter electrode and the measuring electrode can be provided by various methods as shown in FIGS. For example, as shown in FIG. 1, the same plane type provided on the same support, FIG. 2 and FIG.
Modifications such as a multi-layered cylindrical type as shown in FIG. 4 and a three-dimensional type as shown in FIGS. 4 and 5 are possible. When they are provided on the same support, a metal layer of copper, silver, gold, mercury, etc. is formed by vapor deposition, sputtering, etc. on a non-conductive support such as a glass plate, resin plate, resin film, etc. In the conductive substrate and the counter electrode. Further, the organic CT complex formed directly on the conductive substrate at the measurement 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, does not react specifically with proteins, electrolytes, biological components, etc. in the sample liquid, and is used for electrochemical reactions other than enzyme reaction in the measuring electrode. The shape is preferably the same shape or at least the same area for the purpose of compensating for the current value, though it is not particularly limited as long as it is affected by the secondary side reaction and adsorption to the same extent. .. An example of the case where the compensation pole is provided is shown in FIG. 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,
In the step of forming the organic CT complex directly on the conductive substrate in the measuring electrode and the step of applying the water-insoluble polymer, the compensating electrode can be simultaneously prepared, 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.

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 the 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. A comparison between the steady state method and the single pulse method is shown in FIG. 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.

Substances that can be measured by the biosensor of the present invention include sugars such as glucose, trace biological substances in blood and urine such as lactic acid and alcohol, sugars and alcohols in food processing processes. With conventional biosensors, it was necessary to dilute and stir at the time of measurement, but using the biosensor of the present invention, it is possible to measure the sample as it is without diluting and stirring, and it is possible to selectively enhance the above substances. It is possible to perform repeated analysis with high accuracy and over a long period of time.

[0045]

【Example】

[0046]

Example 1 Copper was vacuum-deposited on a glass plate, and a measurement electrode portion was formed.
(1 mm × 1 mm) and a counter electrode portion (10 mm ² ) surrounding it were obtained. Parts other than these electrode parts were molded with an epoxy resin.

1.0 g of 7,7 ′, 8,8′-tetracyanoquinodimethane (reagent, manufactured by Kishida Chemical Co., Ltd., abbreviated as TCNQ hereinafter) was added to 30 ml of acetonitrile (for reagents and spectra) to give a saturated solution of TCNQ. Prepared. The above electrode system was immersed in this TCNQ saturated solution and left for 1 minute. Immediately after the immersion, fine purple fine needle-shaped crystals began to grow on the surface of the electrode part, and after 1 minute, a conductive layer of an organic CT complex of about 2 μm was obtained at the measurement electrode part and the counter electrode part.

2.0% by weight of polyvinyl butyral resin [trade name: S-REC B, BX-L, manufactured by Sekisui Chemical Co., Ltd.] and 2.0% by weight of 1,1′-dimethylferrocene (reagent, manufactured by Tokyo Kasei). Prepare an acetone solution containing 1 ml of glucose oxidase (Aspergillus niger
Origin, Sigma, Type VII) 4.0 mg was added, and 5 μl of a solution that was uniformly dispersed by an ultrasonic disperser was added dropwise to the above-mentioned measurement electrode portion, dried, and then washed with pure water to obtain the measurement electrode. The enzyme was immobilized on.

On the above electrode system, 20 μl of 100 mM phosphate buffer solution (pH 7.0) having each glucose concentration shown in Table 1 was dropped,
The entire surface of the measurement electrode and the counter electrode was covered with the sample solution. The response current for each glucose concentration shown in Table 1 was measured by applying a pulse potential of 0.35 V to the counter electrode against the counter electrode and measuring the current value 5 seconds after the potential was applied. The results are shown in Table 1. Thus, a good linear relationship was obtained in the measured concentration range. Next, the buffer solution of each glucose concentration used above was subjected to a nitrogen bubble for 30 minutes, and then similarly measured.
The results are shown in Table 1. From this result, it can be seen that the response current of the present enzyme electrode is hardly affected by the dissolved oxygen concentration.

Further, 100 mM phosphate buffer (pH 7.0) containing 5 mM ascorbic acid instead of the buffer used above.
Was measured in the same manner. As is clear from the results shown in Table 1, the change in response current value is hardly affected by ascorbic acid. Further, even after the sensor was left in a phosphate buffer for 30 days, the responsiveness of about 80% of the initial value was maintained, and the storage stability was excellent.

[0051]

[Example 2] Instead of the copper vapor-deposited substrate used in Example 1, a silver vapor-deposited plate of the same shape was used and immersed in a TCNQ saturated solution in the same manner as in Example 1 to form a conductive layer composed of an organic CT complex having a thickness of about 1 µm. Layers were obtained. Ferrocenecarboxylic acid (reagent, Aldric
20 μl of an aqueous solution in which 5 mg of glucose oxidase was dissolved in 1 ml of a 5 mM aqueous solution (manufactured by Company H) was dropped only on the measurement electrode part of the silver electrode and dried. After that, a 1% by weight solution of copolymerized polyester resin PES-110L [trade name, manufactured by Toagosei Chemical Industry Co., Ltd.] in methylene chloride as a water-insoluble polymer was applied, dried, and washed with pure water to produce an enzyme. An immobilized electrode was obtained. Table 1 shows the results of measurement using this electrode in the same manner as in Example 1.

[0052]

Example 3 The copper-deposited glass substrate having the measurement electrode portion and the counter electrode portion used in Example 1 was placed in a vacuum vapor deposition apparatus. While the copper plate was heated to about 250 ° C. under a reduced pressure of about 1 × 10 ^-5 mmHg, TCNQ was gradually heated to obtain a conductive layer composed of an organic CT complex having a thickness of about 5 μm in 30 minutes. The glass substrate having the organic CT complex conductive layer was molded with an epoxy resin except for the measurement electrode part and the counter electrode part. Glucose oxidase
20 μl of a 10 mg / ml aqueous solution was dropped on the measurement electrode part of the above electrode and dried. Then, 20 μl of an acetone solution containing 2.0% by weight of 1,1′-dimethylferrocene and 1.0% by weight of polyvinyl butyral resin was applied, dried, and washed with pure water to immobilize the enzyme on the measurement electrode portion. Further, 20 μl of a saturated solution of TCNQ acetonitrile was applied and naturally dried. Table 1 shows the results of measurement using this electrode in the same manner as in Example 1.
Shown in.

[0053]

Example 4 Using the glass substrate having the electrode system composed of the organic CT complex obtained in Example 1, an aqueous solution of potassium octacyanomolybdate was applied and dried. An acetone solution containing 2.0% by weight of polyvinyl butyral resin was prepared, 4.0 mg of glucose oxidase was added to 1 ml of this solution, and 5 μl of the solution uniformly dispersed by an ultrasonic disperser was dropped on the measurement electrode portion of the substrate and dried. Then, it was washed with pure water to immobilize the enzyme on the measurement electrode. Using this electrode, Example 1
The results measured in the same manner as in Table 1 are shown in Table 1.

[0054]

[Reference Example] Using the electrode system obtained in Example 2, a calibration curve was similarly prepared using 0.1 M KCl aqueous solution having glucose concentrations of 2.0 mM and 5.0 mM, and then human serum was used as an unknown sample.
μl was dropped on the electrode system and the response current was measured. The glucose concentration in the serum was determined from the result of the calibration curve prepared in advance, and it was 4.8 mM. As a result of measuring the glucose concentration by a known enzyme colorimetric method, it was 4.8 mM, showing good agreement.

[0055]

[Table 1]

[0056]

[Embodiment 5] Copper is vacuum-deposited on a glass plate, and as shown in FIG.
m) and the surrounding counter electrode part (10 mm ² ) were obtained. Parts other than these electrode parts were molded with an epoxy resin.

1.0 g of 7,7 ', 8,8'-tetracyanoquinodimethane (reagent, manufactured by Kishida Chemical Co., Ltd., abbreviated as TCNQ hereinafter) was added to 30 ml of acetonitrile (reagent, for spectra) to give a saturated solution of TCNQ. Prepared. The above electrode system was immersed in this TCNQ saturated solution and left for 1 minute. Immediately after immersion, fine purple needle-shaped crystals began to grow on the surface of the electrode part, and after 1 minute, about 2 μm of organic material was applied to the measurement pole, compensation pole and counter pole.
A conductive layer composed of CT complex was obtained.

2.0% by weight of polyvinyl butyral resin (trade name: S-REC B, BX-L, manufactured by Sekisui Chemical Co., Ltd.) and 2.0% by weight of 1,1'-dimethylferrocene (reagent, manufactured by Tokyo Kasei). Prepare an acetone solution containing 1 ml of glucose oxidase (Aspergillus niger
Origin, Sigma, Type VII) 4.0 mg was added, and 5 μl of a solution that was uniformly dispersed by an ultrasonic disperser was added dropwise to the above-mentioned measurement electrode portion, dried, and then washed with pure water to obtain the measurement electrode. The enzyme was immobilized on. Next, a solution containing no enzyme in the above solution was prepared in the same manner, and the solution was dropped on the compensating electrode portion in the same manner and dried.

On the above electrode system, 20 μl of 100 mM phosphate buffer solution (pH 7.0) having each glucose concentration shown in Table 2 was dropped,
The entire surface of the measurement electrode, the compensation electrode and the counter electrode was covered with the sample solution. Apply 0.35V pulse potential to the measurement electrode against the counter electrode,
The current value 5 seconds after the application of the potential was measured. Subsequently, a pulse potential having the same potential as that of the counter electrode was applied to the compensation electrode, and a current value 5 seconds after the potential was applied was measured in the same manner, and the difference between the measured current values of the two was taken as a response current value. The results are shown in Table 2. Thus, a good linear relationship was obtained in the measured concentration range.

Next, the buffer solution of each glucose concentration used above was subjected to a nitrogen bubble for 30 minutes and then measured in the same manner.
The results are shown in Table 2. From this result, it can be seen that the response current of the present enzyme electrode is hardly affected by the dissolved oxygen concentration.

Further, 100 mM phosphate buffer (pH 7.0) containing 5 mM ascorbic acid instead of the buffer used above.
Was measured in the same manner. As is clear from the results shown in Table 2, the response current value change is hardly affected by ascorbic acid. Further, even after the sensor was left at room temperature in a desiccator for 30 days, the responsiveness of about 80% of the initial value was maintained, and the storage stability was excellent.

[0062]

[Example 6] Instead of the copper vapor-deposited substrate used in Example 5, a silver vapor-deposited plate of the same shape was used, and by dipping it in a TCNQ saturated solution in the same manner as in Example 5, a conductive layer composed of an organic CT complex having a thickness of about 1 µm Layers were obtained. Ferrocenecarboxylic acid (reagent, Aldric
20 μl of an aqueous solution in which 5 mg of glucose oxidase was dissolved in 1 ml of a 5 mM aqueous solution (manufactured by Company H) was dropped only on the measurement electrode part of the silver electrode and dried. Next, a solution containing no enzyme in the above solution was prepared in the same manner, dropped only on the compensation electrode portion, and dried.

Then, a 1% by weight solution of methylene chloride in a copolyester resin PES-110L [trade name, manufactured by Toagosei Kagaku Kogyo Co., Ltd.] as a water-insoluble polymer was applied to the measuring electrode part and the compensating electrode part and dried. After that, by washing with pure water, an enzyme-immobilized electrode and a compensation electrode were obtained. Table 2 shows the results of measurement using this electrode in the same manner as in Example 5.

[0064]

Example 7 The copper vapor-deposited glass substrate having the measurement electrode portion, the compensation electrode portion and the counter electrode portion used in Example 5 was placed in a vacuum vapor deposition apparatus. Approximately 1 × 10 ^-5 mmHg of the above substrate under reduced pressure
While heating at 200 ° C., TCNQ was gradually heated to obtain a conductive layer composed of an organic CT complex having a thickness of about 5 μm in 30 minutes. The glass substrate having the organic CT complex conductive layer was molded with an epoxy resin except for the measurement electrode part and the counter electrode part. 20 μl of 10 mg / ml glucose oxidase aqueous solution was dropped only on the measurement electrode part of the above electrode and dried. Then, 20 μl of an acetone solution containing 2.0% by weight of 1,1′-dimethylferrocene and 1.0% by weight of polyvinyl butyral resin was applied to the measurement electrode part and the compensation electrode part, dried, and then washed with pure water. Further, 20 μl of a saturated solution of TCNQ acetonitrile was applied and naturally dried. Table 2 shows the results of measurement using this electrode in the same manner as in Example 5.
Shown in.

[0065]

Example 8 Using a glass substrate having an electrode system made of the organic CT complex obtained in Example 5, an aqueous solution of potassium octacyanomolybdate was applied to the entire surface and dried. Prepare an acetone solution of 2.0% by weight of polyvinyl butyral resin,
Glucose oxidase (4.0 mg) was added to 1 ml of this solution, and 5 μl of the solution was uniformly dispersed by an ultrasonic disperser.
Was dropped only on the measurement electrode portion of the substrate and dried, and then a solution containing no enzyme in this solution was dropped only on the compensation electrode portion and dried, and the whole surface was washed with pure water. Table 2 shows the results of measurement using this electrode in the same manner as in Example 5.

[0066]

[Example 9] Using the electrode system obtained in Example 6, and using each solution having a glucose concentration of 4.0 mM, 6.4 mM, 8.2 mM, and 9.6 mM prepared by adding glucose to human serum, The measurement was performed sequentially by the method shown in 5. The result is shown in FIG. 7. Thus, a good linear relationship was obtained even in the measurement using serum.

[0067]

[Table 2]

[0068]

Since the biosensor of the present invention directly detects the electron transfer accompanying the enzymatic reaction, it is not affected by the amount of dissolved oxygen and is not affected by the electrochemical interfering substance. It is also a biosensor with excellent responsiveness and response life. Moreover, since the measuring electrode and the counter electrode are arranged in the vicinity, it is not necessary to dilute and stir a small amount of sample,
It can be used as it is and can be quantified quickly and with high accuracy by simple operation. Further, if a compensation electrode is further provided, it can be quantified with higher accuracy without being affected by side reactions or adsorption of other substances.

[Brief description of drawings]

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

FIG. 2 is a diagram showing an example of the biosensor of the present invention.

FIG. 3 is a diagram showing an example of the biosensor of the present invention.

FIG. 4 is a diagram showing an example of the biosensor of the present invention.

FIG. 5 is a diagram showing an example of the biosensor of the present invention.

FIG. 6 is a diagram showing an example of the biosensor of the present invention.

FIG. 7 is a diagram showing the relationship between glucose concentration and response current measured in Example 9.

FIG. 8 is a comparison diagram measured by the steady state method and the single pulse method.

[Explanation of symbols]

 1: Substrate 2: Counter electrode 3: Measurement electrode 4: Compensation electrode 5: Insulation film 6: Lead part 7: Counter electrode terminal 8: Measurement electrode terminal 9: Compensation electrode terminal

─────────────────────────────────────────────────── ─── Continuation of the front page (51) Int.Cl. ⁵ Identification code Internal reference number FI technical display location 7235-2J G01N 27/46 336 B

Claims (6)

[Claims] 1. A biosensor having an electrode system including a measuring electrode having an enzyme having a conductive layer containing an organic charge transfer complex crystal and a counter electrode provided in the vicinity of the measuring electrode. 2. The measuring electrode has a conductive base and a conductive layer made of an organic charge transfer complex crystal provided on the surface of the base,
The biosensor according to claim 1, which is an electrode on which an enzyme and an electron mediator are immobilized with or without a water-insoluble polymer. 3. The biosensor according to claim 1, further comprising a compensation electrode on which the enzyme is not immobilized. 4. An electrode having a conductive substrate and a conductive layer made of an organic charge transfer complex crystal provided on the surface of the substrate, on which an electron mediator is immobilized with or without a water-insoluble polymer. The biosensor according to claim 2, further comprising a compensation pole. 5. The biosensor according to claim 1, wherein the enzyme is a redox enzyme. 6. A method for analyzing a specific substance by applying a pulse potential to the biosensor according to claim 1 and measuring a response current due to an enzymatic reaction.