JP4185423B2 - Method for producing electrochemical biosensor - Google Patents

Description

  The present invention relates to a biosensor for continuously or repeatedly measuring a biological component contained in a living body or in a biological sample with good quantitativeness.

  In order to elucidate clinical tests and biological functions, measurement of biological components in a living body or a sample sampled from a living body has been actively performed, and many techniques have been reported. Many electrochemical biosensors have been studied in which an oxidase is modified on an electrode. Among them, there are many reports of glucose sensors modified with glucose oxidase. (1) Hydrogen peroxide generated by enzymatic reaction between glucose oxidase and glucose is directly oxidized on the electrode or reduced through horseradish peroxidase. (2) A method for measuring the dissolved oxygen concentration decreased by the enzymatic reaction with glucose oxidase (see Non-patent document 2), (3) Between glucose oxidase and the electrode It can be roughly divided into three methods for directly measuring electron movement (see Non-Patent Document 3).

  In the method of measuring by directly oxidizing the hydrogen peroxide generated by the enzyme reaction of (1) on the electrode, the measurement is usually performed with the electrode to which a potential of about 0.5 V is applied, so ascorbine present in the living body. There is a problem that easily oxidizable substances such as acids act as interfering substances, and the measurement in vivo lacks quantitativeness. Moreover, since oxygen is required for the enzyme reaction, the sensor sensitivity is greatly influenced by the dissolved oxygen concentration.

  Next, the method (2) is not suitable for measurement in the living body because the dissolved oxygen concentration itself in the living body fluctuates. For this reason, as in the method (3), a sensor has been developed that uses an electron transfer mediator to perform direct electron transfer between an enzyme and an electrode and does not require oxygen for the reduction of the enzyme, and performs measurement. This sensor is hardly affected by the dissolved oxygen concentration and can be measured at a relatively low potential. However, in the measurement in the living body, protein is adsorbed on the enzyme membrane surface or the electrode surface, so that the sensitivity of the sensor is lowered and there is a problem that the long-term stability is lacking. Therefore, it is necessary to detect proteins with high quantitativeness by immobilizing various biocompatible membranes on the sensor to inhibit protein adhesion.

  As a compatible membrane for blood, polymers having high hydrophobicity such as polytetrafluoroethylene, polydimethylsiloxane, and polyurethane have been used so that blood components do not adhere to the material surface (see Non-Patent Document 4). Contrary to this, the surface of the material is made of a highly hydrophilic polymer such as poly (2-hydroxyethyl methacrylate) or polyacrylamide to form a surface with low interface free energy, and an attempt to suppress adhesion of blood components is also made. (See Non-Patent Document 4). For biosensors that measure biological components in blood, coating with a hydrophobic membrane inhibits the diffusion of molecules to be measured, making it impossible to measure. Therefore, a hydrophilic biocompatible membrane is mainly fixed to the outermost layer of the biosensor. Have been used (see Non-Patent Document 5). In recent years, sensors having a structure that suppresses adsorption of platelets by electrolytic polymerization of organopolysiloxane with glucose oxidase have also been reported. However, since the electropolymerization takes time, there is a problem that it lacks mass productivity. (See Non-Patent Document 6).

P. G. Osborne, O. Niwa, T. Kato, K. Yamamoto, J. Neurosci. Methods, 77 (1997), 143-150 D. A. Gough, J. Y. Lucisano, P. H. S. Tse, Anal. Chem., 1985, 57, 2351-2357 Y. Degani A. Heller, J. Phys. Chem., 91 (1987), 1285-1289 Supervised by Nobuo Nakabayashi, Development and application of medical polymer materials, ISBN4-88231-202-6 Nishida, Sugawara, Ichinose, Shimoda, Konno, Uemura, Uehara, Shichiri, Ishihara, Nakabayashi, Artificial Organ, 1996, 25, 144 Inoue, Yasawa, Imai, Chemical Sensors, vol.18, Supplement A, 2002, p. 136

  In the measurement of biological samples using an electrochemical biosensor, the sensitivity of the sensor decreased due to protein adsorption, and it was not possible to detect with good quantitativeness. For this reason, various biocompatible membranes have been applied to biosensors in order to suppress sensitivity reduction due to protein adsorption. However, when an organic solvent is used in the formation of a biocompatible film, it is necessary to irradiate ultraviolet rays, and there has been a problem that the activity of the enzyme immobilized on the electrode is reduced during the formation of the biocompatible film. Further, since the biocompatible membrane is formed on the outermost layer of the sensor, diffusion of the target substrate is suppressed, so that there is a problem that the sensitivity of the sensor is lowered and a quick response cannot be obtained.

  In order to solve the above problems, a biosensor was constructed in which a complex of a hydrophobic mediator and a matrix immobilizing an enzyme was formed in an electron transfer mediator and an enzyme formed on an electrode. By combining the hydrophobic mediator and the enzyme membrane, the adsorption of biological components such as proteins is suppressed as well as being unaffected by changes in the oxygen concentration in the sample, and the repeated sensor can be used for measuring biological substances.

  Since the electrochemical biosensor of the present invention has a composite film that prevents protein adsorption and contains an electron transfer mediator, it suppresses the decrease in sensitivity due to protein adsorption, prolongs the sensor life, and affects the influence of oxygen concentration in the sample. It is possible to measure a highly quantitative biological sample that is not received. And since this composite membrane diffuses the target substrate rapidly, high sensitivity and response speed can be obtained.

  Further, in the method for producing an electrochemical biosensor of the present invention, neither an organic solvent nor ultraviolet irradiation is used at the time of application or after application of the enzyme. It is possible to manufacture.

  The electrochemical biosensor of the present invention has a structure having an electrode (working electrode) and a composite film covering the working electrode and containing a matrix, an enzyme, and an electron transfer mediator. The working electrode may be formed of any material known in the art (eg, gold, platinum, silver, carbon, tin oxide, ITO, conductive diamond, organic conductor). The working electrode may have any shape known in the art. For example, the working electrode may have a structure in which a gold wire having a desired cross-sectional area is covered with a polymer such as an epoxy resin so that only the end surface of the gold wire is exposed.

  The composite film is a film including a matrix, an enzyme, and a low-molecular electron transfer mediator having a hydrophobic functional group. The enzyme is capable of transferring electrons to and from the working electrode via an electron transfer mediator, and is capable of oxidizing / reducing a compound to be detected, as long as it is any known in the art. It may be a thing. For example, it can be applied to various biosensors for detecting these substances using enzymes that oxidize / reduce glucose, lactic acid, glutamic acid, alcohol, cholesterol and the like.

  The electron transfer mediator is a component that is confined in the composite membrane to form a conductive electron transfer path between the working electrode and the enzyme. As the electron transfer mediator, it is preferable to use a hydrophobic, low molecular weight organic compound or complex. In the present invention, “hydrophobic” means that the electron transfer mediator dissolved in a solution at a saturated concentration does not give a redox peak current by cyclic voltammetry. Further, “low molecular weight” in the present invention means that the molecular weight of the monomer is 300 or less. Furthermore, the electron transfer mediator used in the present invention preferably has a functional group such as a crosslinkable vinyl group, hydroxyl group, amino group or carboxylic acid. A vinyl group is particularly useful. Preferred electron transfer mediators in the present invention include vinyl ferrocene and its derivatives. Vinyl ferrocene derivatives that can be used include methyl vinyl ferrocene, hydroxy vinyl ferrocene, and amino vinyl ferrocene derivatives.

  Albumin is a matrix that imparts biocompatibility (particularly blood compatibility) to the composite membrane and serves as a carrier for immobilizing the enzyme. As a matrix, polypeptides such as albumin-derived proteins insolubilized by cross-linking reaction, polylysine obtained synthetically, polyamino acids such as polyaspartic acid, other polyacrylamide, polyvinyl alcohol, polyethylene glycol, dextran, carboxymethylcellulose Can be used. Biological matrix materials such as albumin can be easily cross-linked under mild conditions such as room temperature by using a cross-linking agent such as glutaraldehyde. Enzymes and electron transfer mediators can be immobilized. Also, certain enzymes and electron transfer mediators can be cross-linked with the matrix and / or cross-linking agent, so the enzymes and electron transfer mediators may be cross-linked with the matrix by chemical bonds.

  The electrochemical biosensor of the present invention uses a step of applying a solution (dispersion) containing an electron transfer mediator to a working electrode and a step of applying an aqueous solution (or aqueous dispersion) containing a matrix and an enzyme. Can be manufactured.

  The step of applying the electron transfer mediator is a casting method, a dip coating method or a spin method using a solution in which a hydrophobic electron transfer mediator is dissolved in a low boiling point organic solvent such as ethanol or acetone, or an aqueous dispersion of the electron transfer mediator. It can be carried out by any method known in the prior art, such as a coating method. When a low-boiling organic solvent is used, it is desirable to volatilize the solvent before the matrix and enzyme coating step. Thereby, it is possible to suppress a decrease in enzyme activity due to the solvent.

  Similarly, the step of applying the matrix and the enzyme can be performed by any method known in the prior art, such as a casting method, a dip coating method, or a spin coating method. The aqueous solution (aqueous dispersion) containing the matrix and the enzyme may further contain a crosslinking agent such as glutaraldehyde. By applying a solution further containing a cross-linking agent, the cross-linking of the matrix or the cross-linking of the matrix and the enzyme or the electron transfer mediator is advanced simultaneously with the diffusion of the electron transfer mediator into the matrix. A composite membrane can be formed. At this time, the enzyme may be immobilized by being chemically bonded to the matrix via the cross-linking agent, or may be immobilized by being incorporated into the matrix / cross-linking agent network. The same applies to the electron transfer mediator.

  Alternatively, the composite membrane of the present invention may be formed by applying a matrix / enzyme aqueous solution (aqueous dispersion) that does not contain a crosslinking agent, and then allowing the crosslinking agent to act. The cross-linking agent can act in the liquid phase or the gas phase. For example, it is possible to apply a solution containing a cross-linking agent or to be exposed to the vapor of the cross-linking agent. In this case, after the diffusion of the electron transfer mediator into the matrix, the cross-linking of the matrix or the cross-linking of the matrix and the enzyme or the electron transfer mediator proceeds by the action of the cross-linking agent. At this time, the enzyme may be immobilized by being chemically bonded to the matrix via the cross-linking agent, or may be immobilized by being incorporated into the matrix / cross-linking agent network. The same applies to the electron transfer mediator.

  Hereinafter, the present invention will be specifically described. In addition, this invention is not limited only to a following example.

(Example 1)
FIG. 1 shows a schematic diagram of a cross section of an electrochemical glucose sensor structure according to the present invention. The electrochemical glucose sensor of FIG. 1 has a gold electrode 1 having a diameter of 1.6 mm, an epoxy resin 2, a vinyl ferrocene layer 3, and a vinyl ferrocene-albumin-glucose oxidase composite layer 4.

  A 1.6 mm diameter gold electrode (manufactured by BAS) 1 was polished with ion-exchanged water and air-dried after polishing the electrode surface. Thereafter, an ethanol solution (0.1 M) of vinyl ferrocene was cast on a 4 μl electrode and dried at room temperature for 10 minutes to form a vinyl ferrocene layer 3 on the electrode. Furthermore, a solution obtained by mixing glucose oxidase with a 2% bovine serum albumin solution at a weight ratio of 2% and further adding glutaraldehyde to a final concentration of 0.5% is placed on the vinylferrocene layer 3. 2 μl was applied. By applying an albumin solution that has been cross-linked with glutaraldehyde on the vinyl ferrocene layer, the vinyl ferrocene was taken into the albumin-enzyme film to form a vinyl ferrocene-albumin-glucose oxidase composite film 4.

  FIG. 2 shows the result of analyzing the film formed on the electrode using X-ray photoelectron spectroscopy (XPS). FIG. 2 (a) shows the result of elemental analysis by XPS measurement of the substrate surface on which only the vinylferrocene layer 3 is formed on the gold electrode, and FIG. 2 (b) shows the composite film formed on the gold electrode by the method of the present invention. 4 shows the results of elemental analysis of the film surface of No. 4. In the spectrum of FIG. 2A, in addition to iron contained in ferrocene, a gold peak of the electrode is detected. That is, the electrode is not completely covered by the vinylferrocene layer. However, in the composite film 4 of FIG. 2B, it can be confirmed that the analysis of only the film surface is performed without detecting the gold peak of the substrate. In addition to nitrogen and sulfur contained in albumin, iron is detected. This indicates that the vinyl ferrocene-albumin-glucose enzyme complex film is formed by the vinyl ferrocene formed on the gold electrode surface by this method diffusing into the coated albumin film and reaching the vicinity of the film surface. Show. Moreover, iron existing on the surface does not decrease so much even after measurement using a glucose measuring device described later, suggesting that vinylferrocene is stably held in the enzyme-matrix membrane.

  FIG. 3 shows a schematic diagram of a glucose measuring device using a glucose sensor produced by the above method. 3 includes a glucose sensor 5, a silver / silver chloride reference electrode 6, a platinum wire 7, a pH 7 phosphate buffer 8, a stirrer 9, a magnetic stirrer 10, a potentiostat 11, a computer 12, and a glass container 13. Including. The glucose sensor 5 according to the present invention prepared as described above was immersed in the phosphate buffer 8 filled in the glass container 13. Similarly, a silver / silver chloride electrode 6 (manufactured by BAS) as a reference electrode and a platinum wire having a diameter of 1 mm as a counter electrode were immersed in the phosphate buffer 8. The glucose sensor 5, the silver / silver chloride reference electrode 6, and the platinum wire 7 were respectively connected to the working electrode, reference electrode, and counter electrode terminal of the potentiostat 11 (manufactured by CHI). A potential of 0.3 to 0.7 V (vs. silver / silver chloride reference electrode 6) was applied to the glucose sensor 5 using the potentiostat 11, and the change in the current value was recorded and stored by the computer 12. Further, the phosphate buffer 8 was stirred at a constant speed by the stirring bar 9 and the magnetic stirrer 10 placed in the glass container 14.

  FIG. 4A shows a change in response current value of the glucose sensor 5 according to the present invention. A potential of 0.3 V is applied to the electrode (working electrode) of the glucose sensor 5 with respect to the silver / silver chloride reference electrode 6 (reference electrode) using the potentiostat 11 to obtain a stable current value. The glucose solution was added so that the concentration was 5 mM. Immediately after the glucose solution injection, the oxidation current value increased and showed a constant current value of about 0.3 μA. Furthermore, horse serum was injected into the phosphate buffer so that the final concentration was 10%, but the current value did not change at all. From this, it was confirmed that the glucose sensor of the present invention can detect glucose with high quantitativeness even in the presence of a protein contained in serum.

(Comparative Example 1)
A hydrogen peroxide direct oxidation glucose sensor using a platinum electrode was prepared, and glucose measurement was performed in the presence of horse serum.

  The hydrogen peroxide direct oxidation glucose sensor at the platinum electrode is a 1.6 mm diameter platinum electrode mixed with 2% bovine serum albumin at 2% by weight glucose oxidase, and a final concentration of 0.5%. Then, 4 μl of a solution to which glutaraldehyde was added was applied and laminated on the platinum electrode to obtain a glucose sensor.

  FIG. 4B shows a change in current value when glucose is measured using a hydrogen peroxide direct oxidation glucose sensor. Using the same measurement apparatus as in Example 1, a potential of 0.5 V (vs. reference electrode) was applied to the platinum electrode of the sensor to obtain a stable current value, and the final concentration became 5 mM. Glucose solution was added as follows. After the solution injection, the oxidation current value increased and showed a constant current value of about 0.8 μA.

  In this state, when horse serum was added to the buffer solution in the measuring device so that the final concentration was 10%, the oxidation current value decreased, and the oxidation current value was about half and showed a constant value (Fig. 4 (b)). The above results show that in the measurement of glucose by direct oxidation of hydrogen peroxide using a platinum electrode, serum proteins adhere to the matrix layer immobilized on the electrode in the presence of serum, and glucose and hydrogen peroxide diffuse. This shows the conventional problem that the sensor sensitivity is greatly reduced due to the hindrance.

(Comparative Example 2)
A glucose sensor using an osmium polyvinylpyridine complex (Os-gel-HRP) as an electron transfer mediator was prepared, and glucose was measured in the presence of horse serum.
A glucose sensor using Os-gel-HRP as a mediator was prepared as follows as a biosensor for detecting hydrogen peroxide produced by an enzymatic reaction by reduction. First, 2 μl of Os-gel-HRP was modified on a platinum electrode having a diameter of 1.6 mm by a casting method. After drying at room temperature for 1 hour, glucose oxidase was mixed with 2% bovine serum albumin so that the weight ratio was 2%, and glutaraldehyde was added to a final concentration of 0.5%. 2 μl of -gel-HRP was coated on a modified platinum electrode and laminated.

  FIG. 4C shows a change in current value when glucose is measured using a glucose sensor using Os-gel-HRP as an electron transfer mediator. Using the same measuring apparatus as in Example 1, a potential of 0 V was applied to the sensor electrode (working electrode) with respect to the reference electrode to obtain a stable current value, and the final concentration was 5 mM. Glucose solution was added. After the solution injection, the reduction current value increased and showed a constant current value of about 0.5 μA.

  In this state, when horse serum was added to the buffer solution in the measuring device so that the final concentration was 10%, the reduction current value decreased, and the reduction current value was about half and showed a constant value (FIG. 4 ( c)). This shows the conventional problem that the protein in serum adheres to the matrix layer fixed on the electrode, the diffusion of glucose and hydrogen peroxide is inhibited, and the sensitivity is lowered.

  The results of these comparative examples also show that the glucose sensor according to the present invention inhibits protein adhesion by the vinylferrocene-albumin-glucose oxidase complex layer even in the presence of serum and can detect glucose with good quantitativeness in serum. ing.

(Example 2)
A vinylferrocene layer was formed on a gold electrode having a diameter of 1.6 mm in the same manner as in Example 1. Thereafter, 4 μl of a solution in which lactate oxidase was mixed with 2% bovine serum albumin to a weight ratio of 2% and glutaraldehyde was added to a final concentration of 0.5 was applied onto vinylferrocene.

  Using a device similar to that of Example 1, a potential of 0.3 V (vs. silver / silver chloride reference electrode) was applied to the electrode of the lactic acid sensor produced by the above method to obtain a stable baseline. Lactic acid was added so that the final concentration was 5 mM. Thereafter, the oxidation current value gradually increased and showed a constant current value at about 0.2 μA.

  The above results indicate that not only glucose sensors but also lactic acid, glutamic acid, alcohol, cholesterol, etc. can be obtained by immobilizing an enzyme that reacts with a target substrate on an enzyme laminated on a hydrophobic mediator immobilized on an electrode. It can be applied to various biosensors.

(Example 3)
Similarly to Example 1, a vinyl ferrocene layer 3 was formed on the electrode. Further, 2 μl of a 2% bovine serum albumin solution mixed with glucose oxidase at a weight ratio of 2% was applied onto the vinylferrocene layer 3 and air-dried. This electrode was placed in a container placed on the bottom of 1 ml of 25% glutaraldehyde (manufactured by Sigma), and the mixed membrane of bovine serum albumin and glucose oxidase was insolubilized by the vapor of glutaraldehyde, and the vinylferrocene-albumin-glucose oxidase composite layer 4 Formed.

  As in Example 1, this electrode was connected to a device for measuring glucose. First, cyclic voltammetry was performed in a phosphate buffer not containing glucose, and the potential operation was repeated until the current response of the mediator was stabilized. Thereafter, glucose was measured in the same manner as in Example 1. As a result, glucose could be quantified, and a decrease in sensor sensitivity due to injection of horse serum could be prevented. In this case, the vinyl ferrocene layer 3 is in a solid state. Comparing the cyclic voltammogram before and after immersing this sensor in a buffer solution and adding 50% ethyl alcohol to this solution. The vinyl ferrocene retained on the enzyme-matrix membrane by the hydrophobic interaction was dissolved by adding ethyl alcohol, and the concentration of vinyl ferrocene in the solution increased to give a large current. From this, it was shown that hydrophobic vinylferrocene does not dissolve in the buffer due to its hydrophobicity and exists as a solid state, but functions as a mediator.

It is a typical sectional view of an electrochemical glucose sensor. It is an X-ray electron spectroscopic spectrum of a film formed on a gold electrode, (a) is a spectrum of a film formed only from vinylferrocene, and (b) is a spectrum of the composite film of the present invention. It is a schematic diagram which shows the glucose measuring apparatus using the electrochemical glucose sensor of this invention. It is a graph which shows the response current value of a glucose sensor, (a) shows the response current value of the electrochemical glucose sensor of this invention, (b) shows the response current value of a hydrogen peroxide direct oxidation glucose sensor, c) is a graph showing a response current value of a glucose sensor using Os-gel-HRP as an electron transfer mediator.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 Platinum electrode 2 Epokin resin 3 Vinylferrocene layer 4 Vinylferrocene-albumin-glucose oxidase composite film 5 Glucose sensor 6 Silver / silver chloride reference electrode 7 Platinum wire 8 pH7 phosphate buffer 9 Stirrer 10 Magnetic stirrer 11 Potentiostat 12 Computer 13 Glass container

Claims (4)

Applying an electron transfer mediator on the electrode;
Coating an albumin aqueous solution in which an enzyme and glutaraldehyde are dissolved on the electron transfer mediator to diffuse the electron transfer mediator into albumin;
A method for producing an electrochemical biosensor, comprising: reacting and crosslinking the albumin, an enzyme, and an electron transfer mediator to form a composite film. The method for producing an electrochemical biosensor according to claim 1 , wherein the step of applying the electron transfer mediator is performed by a method selected from the group consisting of a casting method, a dip coating method, and a spin coating method. Applying a hydrophobic electron transfer mediator of vinylferrocene and its derivatives on the electrode;
Coating an albumin aqueous solution in which an enzyme and glutaraldehyde are dissolved on the electron transfer mediator to diffuse the electron transfer mediator into albumin;
A method for producing an electrochemical biosensor, comprising: A step of applying the electron transfer mediator, casting, method for producing an electrochemical biosensor according to claim 3, characterized in that carried out by a method selected from the group consisting of dip coating and spin coating.

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