Biosensor and stirred reaction vessel

Abstract

[Problem] To realize a biosensor (1) which does not require output calibration and which can be used under a wide range of conditions. [Solution] A measuring space (13) is formed in an insertion portion (10) to be inserted into a liquid sample, and when the insertion portion is inserted into the liquid sample, the measuring space is filled with the liquid sample. Furthermore, an electrode (14) is attached within the measuring space, and an enzyme (14z) is fixed onto the electrode in a state of contact with the liquid sample filling the measuring space. In this way, since it is possible to realize a biosensor that does not have a limited transmission layer, the risk of deterioration or damage to the limited transmission layer is eliminated. In addition, since the electrode is attached within the measuring space of the insertion portion, the stationary liquid specimen filling the measuring space can be brought into contact with the electrode. As a result, it is possible to realize a biosensor which does not require calibration of the sensor output and which can be used under a wide range of conditions.

Description

Biosensors, stirring reaction vessels

[0001] The present invention relates to a biosensor that converts a target component contained in a liquid sample into another component using an enzyme immobilized on an electrode, and measures the concentration of the target component by detecting the current value generated at the electrode as a result of the conversion, and a stirring reaction vessel equipped with the biosensor.

[0002] A biosensor is known that measures the concentration of a target component in a liquid sample by converting the target component, such as a protein, contained in the liquid sample into another component using an enzyme immobilized on an electrode and detecting the current value of the current generated at the electrode as a result of this conversion.

[0003] This biosensor generally has a structure in which an enzyme layer is formed on an electrode, and a limiting permeation layer is formed on the enzyme layer. Enzymes are immobilized on the enzyme layer, and the enzymes convert the target components to other components. The limiting permeation layer protects the enzyme layer and also has the function of selectively allowing the target components in the liquid sample to pass through and supply them to the enzyme layer. When a liquid sample is dropped onto the limiting permeation layer, the target components in the liquid sample diffuse through the limiting permeation layer, reach the enzyme layer, and are converted to other components by the enzymes in the enzyme layer. At this time, electrons are exchanged between the enzyme and the target components, generating a current in the electrode.

[0004] Here, the enzyme can convert the analyte component at a sufficient rate, and the rate-limiting step in the reaction in which the enzyme converts the analyte component is the process in which the analyte component is supplied to the enzyme, and therefore the process in which the analyte component diffuses through the limited permeation layer. The rate at which the analyte component diffuses through the limited permeation layer varies depending on the concentration of the analyte component in the liquid sample in contact with the limited permeation layer. Ultimately, the current value generated at the electrode varies depending on the concentration of the analyte component in the liquid sample. As a result, the concentration of the analyte component in the liquid sample can be measured by detecting the current value generated at the electrode (Patent Document 1).

[0005] JP 2000-081409 A (paragraph 0003)

[0006] However, the conventional biosensors described above have the problem that the sensor output needs to be calibrated periodically, and in addition, there are significant limitations on the conditions of use. The reason for this is that when the biosensor is in use, the limiting permeation layer becomes contaminated by various components in the liquid sample, which changes the rate at which the components to be measured diffuse through the limiting permeation layer, resulting in changes in the current value generated at the electrode. Furthermore, there is a risk that the limiting permeation layer will peel off from the enzyme layer due to insufficient strength or swelling of the limiting permeation layer. Therefore, in order to prevent peeling, it is necessary to limit the conditions of use of the biosensor.

[0007] This invention has been made to solve the above-mentioned problems of conventional biosensors, and aims to realize a biosensor that does not require calibration of the sensor output and can be used under a wide range of conditions.

[0008] In order to solve the above-mentioned problems, the biosensor of the present invention employs the following configuration: A biosensor that converts a predetermined target component contained in a liquid sample into a component different from the target component by an enzyme immobilized on an electrode, and measures the concentration of the target component by detecting the value of a current generated at the electrode accompanying the conversion, the biosensor comprising an insertion part to be inserted into the liquid sample, the insertion part forming a measurement space that is filled with the liquid sample when the insertion part is inserted into the liquid sample and the liquid sample flows into the insertion part, the electrode being attached within the measurement space, and the enzyme being immobilized on the electrode in a state of contact with the liquid sample that fills the measurement space.

[0009] Although the detailed reason will be described later, it has been found that, as long as the liquid sample in contact with the biosensor can be kept stationary and free of flow, a current proportional to the concentration of the component to be measured in the liquid sample will be generated at the electrode even if the biosensor does not have a limiting permeation layer. Therefore, in the biosensor of the present invention, a measurement space is formed in the insertion part that is inserted into the liquid sample, and the measurement space is filled with the liquid sample when the insertion part is inserted into the liquid sample. An electrode is then attached to the measurement space, and an enzyme is immobilized on the electrode in a state in which it comes into contact with the liquid sample when the measurement space is filled with the liquid sample.

[0010] This allows for the realization of a biosensor without a limiting permeation layer, thereby eliminating the possibility of contamination or peeling of the limiting permeation layer. Additionally, because the electrodes are attached within the measurement space of the insertion part, even if a flow exists in the liquid sample into which the insertion part is inserted, the influence of that flow within the measurement space can be suppressed. Therefore, a current proportional to the concentration of the component to be measured in the liquid sample is generated in the electrode. As a result, it is possible to realize a biosensor that does not require calibration of the sensor output and can be used under a wide range of conditions.

[0011] In the biosensor of the present invention described above, the measurement space may be formed in the shape of a column with one end open and a bottom, and the electrode may be attached to the bottom surface of the column.

[0012] This allows the measurement space to have a simple structure, making it easy to manufacture the biosensor. Furthermore, the electrode can be attached at a position away from the position where the measurement space opens to the outside. Therefore, even if there is a flow in the liquid sample outside the measurement space, the flow is less likely to affect the liquid sample around the electrode. As a result, it is possible to prevent fluctuations in the output of the biosensor due to the flow in the liquid sample.

[0013] In addition, in the biosensor of the present invention in which the measurement space is formed in the shape of a column with a bottom, the measurement space may be formed by erecting a protruding wall from the periphery of the electrode and surrounding the space around the electrode with the protruding wall.

[0014] Even in this case, the measurement space can have a simple structure, making it easy to manufacture the biosensor. In addition, the electrodes can be positioned away from the position where the measurement space opens to the outside, which makes it possible to prevent fluctuations in the output of the biosensor due to the influence of flow in the liquid sample.

[0015] Furthermore, in the biosensor of the present invention in which the measurement space is formed in a columnar shape with a bottom, the height of the measurement space may be equal to or greater than the diameter of a circle having the same area as the area of ​​the bottom of the columnar shape.

[0016] This allows the distance from the position where the measurement space opens to the outside to be sufficient to the electrode, making it possible to prevent fluctuations in the output of the biosensor due to the influence of flow in the liquid sample.

[0017] Furthermore, the biosensor of the present invention described above is particularly suitable for use in monitoring or managing the progress of various reactions, and therefore the present invention can also be understood as a stirring reaction vessel equipped with a biosensor.

[0018] In the stirring reaction vessel of the present invention, the concentration of the component to be measured can be detected stably and accurately even while the liquid sample is being stirred, so that the progress of the reaction can be monitored accurately and in real time, thereby enabling appropriate management of various reactions.

[0019] 1 is a perspective view showing the external shape of the biosensor 1 of the first embodiment. FIG. 1 is an explanatory diagram showing how an enzyme 14z immobilized on an electrode 14 converts a measurement target component a in a liquid sample into another component b. FIG. 2 is an explanatory diagram showing how a spherical diffusion region 14s is formed around the electrode 14 in a liquid sample. FIG. 3 is an explanatory diagram of an experimental method for investigating the effect of the presence or absence of a flow of a liquid sample on sensor output using a comparative biosensor 9 that does not have a measurement space 13. FIG. 4 is an explanatory diagram showing the results of an experiment investigating the effect of the presence or absence of a flow of a liquid sample on sensor output using the comparative biosensor 9. FIG. 5 is an explanatory diagram for intuitively explaining why the sensor output of the comparative biosensor 9 varies depending on the flow of a liquid sample. FIG. 6 is an explanatory diagram showing the interior of the measurement space 13 by taking a cross section of the tip portion of the biosensor 1 of the first embodiment. FIG. 7 is an explanatory diagram of an experimental method for investigating the effect of the presence or absence of a flow of a liquid sample on sensor output using the biosensor 1 of the first embodiment. FIG. 8 is an explanatory diagram showing the results of an experiment investigating the effect of the presence or absence of a flow of a liquid sample on sensor output using the biosensor 1 of the first embodiment. FIG. 1 is an explanatory diagram illustrating a first modified example of the biosensor 1 of the first embodiment. FIG. 2 is an explanatory diagram illustrating a second modified example of the biosensor 1 of the first embodiment. FIG. 3 is an explanatory diagram illustrating a third modified example of the biosensor 1 of the first embodiment. FIG. 4 is an explanatory diagram about the biosensor 1 of the second embodiment. FIG. 5 is an explanatory diagram for intuitively explaining why the biosensor 1 of the second embodiment is less susceptible to the influence of the flow of a liquid sample. FIG. 6 is an explanatory diagram about a modified example of the biosensor 1 of the third embodiment. FIG. 7 is an explanatory diagram for intuitively explaining why the biosensor 1 of the third embodiment is less susceptible to the influence of the flow of a liquid sample. FIG. 8 is an explanatory diagram illustrating an agitation reaction vessel equipped with the biosensor 1.

[0020] A. First Embodiment FIG. 1 is a perspective view showing the external shape of a biosensor 1 of a first embodiment. As shown in the figure, the biosensor 1 of the first embodiment has a cylindrical insertion section 10 with a flange 11 formed at the base end thereof, and a short cylindrical protruding wall 12 protruding from the tip of the insertion section 10. A cylindrical measurement space 13 with one open end is formed inside the protruding wall 12, and a small electrode 14 is attached to the center of the bottom surface 13a of the measurement space 13. The electrode 14 is a minute electrode with a diameter (or the major axis of the ellipse) of at most 0.1 mm, and the measurement space 13 is also a small space with a cross-sectional diameter of at most 5 mm. A thin gold electric wire (not shown) extends from the electrode 14 and is connected to a lead wire 15 inside the insertion section 10. The lead wire 15 is then extended to the outside from the end face on the side where the flange 11 is formed.

[0021] The size of the electrodes 14 and the measurement space 13 can be appropriately changed depending on the environment in which the biosensor 1 is used (e.g., the flow rate and viscosity of the liquid to be measured) so that the hydrodynamic size (i.e., the Reynolds number) is approximately the same. For example, if the flow rate of the liquid to be measured is low or the viscosity of the liquid is high, the electrodes 14 and the measurement space 13 may be made larger. For example, the viscosity of the liquid may be about 500 times that of water. In such a case, the diameter of the electrodes 14 can be changed to about 50 mm, and the cross-sectional diameter of the measurement space 13 can also be changed to about 50 mm.

[0022] The biosensor 1 of the first embodiment is used by inserting the tip of the insertion portion 10 into a liquid sample. When the tip of the insertion portion 10 is inserted into the liquid sample, the liquid sample flows in from the open end, filling the measurement space 13 with the liquid sample. As a result, the electrode 14 is exposed to the liquid sample. Note that the open end of the measurement space 13 is the portion through which the measurement space 13 communicates with the outside, and is therefore sometimes referred to as the "communication port" below. An enzyme is immobilized on the electrode 14, and the enzyme converts the target component in the liquid sample into another component. During this process, electrons are exchanged between the enzyme and the target component, generating a current in the electrode 14. Note that the enzyme immobilized on the electrode 14 is preferably a type of enzyme that can directly exchange electrons with the electrode 14 without the intervention of an intermediate substance called a mediator (direct electron transfer enzyme). Using this type of enzyme allows the electron exchange between the enzyme and the target component to be directly transmitted to the electrode 14. Furthermore, such enzymes are preferably enzymes that selectively donate and accept electrons to and from oxygen, glucose, lactic acid, amino acids (particularly glutamic acid), ammonia, carbon dioxide, etc. Examples of such enzymes include the following: fructose dehydrogenase, multicopper oxidase enzyme, glucose dehydrogenase, glucose oxidase enzyme, pyranose dehydrogenase, ferredoxin-NADP+ reductase enzyme, cellobiose dehydrogenase, gluconate dehydrogenase, succinate dehydrogenase, histamine dehydrogenase, lactate dehydrogenase, alcohol dehydrogenase, aldehyde dehydrogenase, cytochrome P450 enzyme, peroxidase enzyme, xanthine dehydrogenase, xanthine oxidase enzyme, sulfite oxidase enzyme, sulfite dehydrogenase, formate dehydrogenase, nitrogenase enzyme, nitrate reductase enzyme, carbon monoxide dehydrogenase, hydrogenase enzyme, galactose oxidase enzyme, and nitrous oxide reductase enzyme.

[0023] 2 is an explanatory diagram showing how an enzyme 14z immobilized on an electrode 14 converts a target component in a liquid sample into another component. In FIG. 2, target component a in the liquid sample is converted into another component b by the enzyme 14z. In the example shown in FIG. 2, electrons are extracted from target component a through conversion and flow into the electrode 14. Of course, electrons may also be supplied to target component a through conversion, in which case the electrons are supplied from the electrode 14. A thin gold electric wire 15a is connected to the electrode 14, and the lead wire 15 (see FIG. 1) is connected to the electric wire 15a.

[0024] In a typical biosensor, a limited-permeation layer is formed on an enzyme immobilized on an electrode, and the target component in the liquid sample diffuses through the limited-permeation layer and is then converted by the enzyme. The rate at which the target component diffuses through the limited-permeation layer is proportional to the concentration of the target component in the liquid sample (hereinafter sometimes referred to as the "substrate concentration of the target component"). Therefore, the concentration of the target component determines the rate at which the target component is supplied to the enzyme, and as a result, the current value generated at the electrode. Therefore, the concentration of the target component in the liquid sample is measured by detecting the current value at the electrode.

[0025] In contrast, the biosensor 1 of the first embodiment does not have a limiting permeation layer, and the enzyme 14z on the electrode 14 is directly exposed to the liquid sample (see FIG. 2). The biosensor 1 of the first embodiment is able to achieve this (eliminate the need for a limiting permeation layer) because the enzyme 14z immobilized on the electrode 14 can convert the target component a at a sufficient rate, thereby forming a spherical region in the liquid sample surrounding the electrode 14, where the concentration of the target component a decreases toward the electrode 14 and reaches 0% at the electrode 14. The reason for this will be explained below. Because the target component a moves toward the electrode 14 within this spherical region by diffusion, this region will be referred to as the "diffusion region 14s" below.

[0026] FIG. 3 is an explanatory diagram showing the formation of a spherical diffusion region 14s around the electrode 14 in a liquid sample. As described above, the enzyme 14z on the electrode 14 can convert the analyte component a at a sufficient rate, so the concentration of the analyte component a in the liquid sample at the position of the electrode 14 is 0%. Then, the analyte component a is supplied to this 0% concentration region by molecular diffusion from adjacent regions. Accordingly, the concentration of the analyte component a in the region where the analyte component a was supplied decreases, and the analyte component a is supplied from the outer regions by molecular diffusion. Accordingly, the concentration of the analyte component a also decreases in the outer regions, and the analyte component a is supplied from the outer regions. Furthermore, the decrease in the concentration of the analyte component a due to the supply of the analyte component a to the inner region becomes smaller the further outward the region. Therefore, a spherical diffusion region 14s is formed in the liquid sample, centered on the electrode 14, in which the concentration of the analyte component a varies from 0% to the substrate concentration c.

[0027] When a diffusion region 14s is formed with the electrode 14 at its center, the rate at which the target component a in the liquid sample is supplied to the electrode 14 depends on the rate at which the target component a undergoes molecular diffusion through the diffusion region 14s. As described above, in a typical biosensor, the rate at which the target component diffuses through the limited-permeation layer determines the rate at which the target component is supplied to the enzyme on the electrode. In contrast, in the biosensor 1 of the first embodiment, the rate at which the target component a in the liquid sample undergoes molecular diffusion through the diffusion region 14s determines the rate at which the target component a is supplied to the enzyme 14z. That is, in the biosensor 1 of the first embodiment, the limited-permeation layer, which was essential in a typical conventional biosensor, is replaced with the diffusion region 14s. Because the limited-permeation layer is susceptible to deterioration or damage, the output of the biosensor must be calibrated during use, and the conditions of use are also subject to significant limitations. Therefore, replacing the limited-permeation layer with the diffusion region 14s eliminates the need for calibration and also eliminates the need for significant limitations on the conditions of use.

[0028] Furthermore, when a spherical diffusion region 14s such as that shown in FIG. 3 is formed, it has been found that the rate v of molecular diffusion of the target component a in the liquid sample through the diffusion region 14s to the electrode 14 can be calculated by the following equation: v = 4·Kd·c / (π·r) where Kd is the diffusion coefficient of the target component a, c is the substrate concentration of the target component a, and r is the radius of the electrode 14. Therefore, if the number of electrons generated by converting 1 mol of the target component a is n and the Faraday constant is F, the limiting current density j generated by the conversion of the target component a by the enzyme 14z when the spherical diffusion region 14s such as that shown in FIG. 3 is formed can be calculated by the following equation: j = 4·n·F·Kd·c / (π·r) where Faraday's constant F is a constant value, the number of electrons n is a physical property of the target component a, and the diffusion coefficient Kd is a physical property determined by the combination of the liquid sample and the target component a. Therefore, once the combination of the liquid sample and the target component a and the radius r of the electrode 14 are determined, the limiting current density j will be proportional to the substrate concentration c of the target component a, and a biosensor 1 that does not require calibration can be realized in principle.

[0029] However, the formation of a spherical diffusion region 14s as shown in Figure 3 requires that the liquid sample around the electrode 14 is stationary. Therefore, in a situation where there is a flow in the liquid sample around the electrode 14, the limiting current density j may not be proportional to the substrate concentration c of the component a to be measured. Therefore, the following experiment was conducted to confirm this.

[0030] Figure 4 is an explanatory diagram showing an experimental method for investigating the effect of the presence or absence of a flow of a liquid sample on the output of a biosensor. As shown in Figure 4, a liquid sample and a rotor 21 are placed in a container 20 such as a beaker, and the container 20 is placed on a stirrer 22. Furthermore, a comparative biosensor 9, which does not have a measurement space 13, is prepared in comparison with the biosensor 1 of the first embodiment shown in Figure 1. The tip of the comparative biosensor 9 is then inserted into the liquid sample, and the output of the biosensor 9 is measured while stirring the liquid sample with the stirrer 22 and stopping the stirring.

[0031] Figure 5 is an explanatory diagram showing the experimental results. In the example shown in Figure 5, the biosensor 9 was inserted into the liquid sample with the stirrer 22 stopped. As shown in Figure 5, the sensor output was large immediately after the biosensor 9 was inserted into the liquid sample, but the sensor output rapidly decreased over time and reached a nearly stable value after about 15 seconds. This change in sensor output is thought to correspond to the process in which the electrode 14 of the biosensor 9 converts the target component a in the liquid sample, thereby forming a spherical diffusion region 14s around it.

[0032] When approximately 15 seconds have passed since the insertion of biosensor 9, rotor 21 is rotated using stirrer 22, which first stirs the liquid sample at the bottom of container 20, and this movement is transmitted to the liquid sample at the top of container 20, which also begins to stir. As a result, the sensor output increases sharply with a short delay from the start of stirring by stirrer 22, and the output continues to fluctuate even after the increase. The sensor output in the increased state is approximately 1.5 times that in the stable state before stirring began.

[0033] When the stirrer 22 is stopped after about 25 seconds of stirring (about 40 seconds after the insertion of the biosensor 9), the sensor output drops sharply and stabilizes at approximately the same value as the output before stirring began after about 30 seconds. Note that while the output stabilized in about 15 seconds when the biosensor 9 was inserted into the liquid sample, it took about twice as long for the output to stabilize after stirring with the stirrer 22 was stopped. This is thought to be because the liquid sample in the container 20 continues to rotate for a while even after the stirrer 22 is stopped.

[0034] After the stirrer 22 is stopped, stirring is restarted once the sensor output has stabilized (approximately 40 seconds after the insertion of the biosensor 9). As with the start of stirring described above, the sensor output increases sharply with a short delay, and then fluctuates while maintaining the increased output. When the stirrer 22 is then stopped, the sensor output decreases sharply and stabilizes in approximately 30 seconds, as with the stop described above.

[0035] As described above, the output of biosensor 9 is easily affected by the presence or absence of a flow in the liquid sample. In the above-mentioned experiment, the sensor output increased by approximately 1.5 times when a flow was generated in the liquid sample. Therefore, it was found that the output of biosensor 9 cannot be said to be proportional to the substrate concentration c of the analyte component a in a flowing liquid sample. This is thought to be because the flow of the liquid sample disrupts the spherical diffusion region 14s. That is, in a still liquid sample, the liquid sample outside the diffusion region 14s has a substrate concentration c of the analyte component a, while inside the diffusion region 14s, the concentration of the analyte component a decreases to 0 as it approaches the central electrode 14. However, when the liquid sample is flowing, the liquid sample with a substrate concentration c outside the diffusion region 14s penetrates into the diffusion region 14s and approaches the electrode 14, which is thought to accelerate the diffusion of the analyte component a to the electrode 14 in that area.

[0036] This phenomenon can be intuitively understood by considering the spherical diffusion region 14s as deforming due to the flow in the liquid sample, as shown in Figure 6. That is, in a liquid sample without flow, a spherical diffusion region 14s is formed as shown by the dashed line in the figure. However, when flow is present in the liquid sample, the spherical diffusion region 14s deforms to the shape shown by the dashed line. Hereinafter, the deformed region of the diffusion region 14s is referred to as the deformed diffusion region 14d. Even in the deformed diffusion region 14d, the liquid sample outside the diffusion region 14d has a liquid sample with a concentration of the target component a at a substrate concentration c. However, inside the diffusion region 14d, the target component a gradually decreases to a concentration of 0 as it approaches the electrode 14. Furthermore, as shown by the dashed arrow in Figure 6, the diffusion of the target component a toward the electrode 14 is accelerated in the portion of the deformed diffusion region 14d that extends inside the spherical diffusion region 14s. Furthermore, because the shape of the deformed diffusion region 14d is constantly changing, the rate at which the measurement target component a is diffused to the electrode 14 also changes constantly. This is thought to be the reason why the output of the biosensor 9 during stirring in the experimental results shown in Figure 5 constantly fluctuates.

[0037] Based on the above considerations, in the biosensor 1 of the first embodiment, a measurement space 13 is formed at the tip of the insertion part 10, and an electrode 14 is attached within the measurement space 13 (see FIG. 1). Note that although FIG. 1 shows the measurement space 13 being formed at the tip of the insertion part 10, it may also be formed on the side surface of the insertion part 10. Furthermore, although the shape of the measurement space 13 is cylindrical, it is not limited to a cylindrical shape and may be, for example, a prismatic shape.

[0038] 7 is an explanatory diagram showing the inside of the measurement space 13 by taking a cross section of the tip portion of the biosensor 1 of the first embodiment. A cylindrical protruding wall 12 stands upright from the tip of the insertion part 10, and the measurement space 13 is formed inside the protruding wall 12. The measurement space 13 is a cylindrical space with one end closed and the other end open, and an electrode 14 is attached to the center of the bottom surface 13a, which is the end face of the closed side. Therefore, when the tip of the biosensor 1 of the first embodiment is inserted into a liquid sample, the measurement space 13 is filled with the liquid sample, and a spherical diffusion region 14s is formed around the electrode 14.

[0039] Furthermore, the diameter D of the open end face of the measurement space 13 (i.e., the communication port) is set to a dimension such that the area S of the communication port is larger than the surface area Sd of the spherical diffusion region 14s. The reason for this is that if the area S of the communication port is smaller than the surface area Sd of the spherical diffusion region 14s, the diffusion rate to the electrode 14 will be determined by the diffusion at the communication port, and the current value generated at the electrode 14 will be small. Here, if the diameter of the spherical diffusion region 14s is Ds, the surface area Sd of the diffusion region 14s can be calculated as follows: Sd = 4π ((Ds / 2)**2) / 2. Note that "**2" represents a square. Therefore, the diameter D of the communication port must satisfy the following condition: π·((D / 2)**2) > 2π·((Ds / 2)**2), and as a result, the diameter D of the communication port must be set to a dimension larger than the root of twice the diameter Ds of the spherical diffusion region 14s. The height H of the measurement space 13 is set to a dimension equal to or larger than the diameter D. The reason for this will be described later. It has been found from experience that, under the conditions of this embodiment, the diameter D of the communication port is preferably in the range of 1.5 μm to 3 cm. The area S of the communication port is preferably 3.5 μm, including when the shape of the communication port is not circular. ２ ~5cm ２ It has been found from experience that the range of

[0040] Furthermore, simulation calculations have shown that, assuming the diameter of the electrode 14 is De, the diameter Ds of the diffusion region 14s is approximately 1.6 times De. This calculation result is consistent with our experience from various experiments. Incidentally, the concentration of the target component a increases slowly with increasing distance from the electrode 14, approaching the substrate concentration c, so the diffusion region 14s does not have a clear boundary. Therefore, the diameter Ds of the diffusion region 14s is calculated using the following method. First, a simulation calculation is used to determine the gradient (concentration gradient) of the increase in the concentration of the target component a on the surface of the electrode 14. Next, the distance required for the target component a to reach the substrate concentration c when increasing from its concentration (0%) on the surface of the electrode 14 at that concentration gradient is determined. This distance corresponds to the radius of the diffusion region 14s, and the diameter Ds of the diffusion region 14s is calculated.

[0041] Therefore, assuming Ds = 1.6·De, substituting this into the above-mentioned formula for calculating the surface area Sd and rearranging it, we obtain Sd = 5.12·Se. Here, Se is the surface area of ​​the electrode 14 (= π·(De / 2)**2). As described above, the opening area S of the communication port needs to be at least larger than the surface area Sd of the diffusion region 14s. The surface area Sd is approximately five times the surface area Se of the electrode 14. Therefore, it is considered sufficient for the opening area S of the communication port to be approximately five times the surface area Se of the electrode 14 (in terms of the diameter D of the communication port, approximately 2.3 times the diameter De of the electrode 14). Alternatively, considering that the diameter Ds of the diffusion region 14s is estimated to be small, it is considered sufficient for the opening area S of the communication port to be approximately 10 times the surface area Se of the electrode 14 (in terms of the diameter D of the communication port, approximately five times the diameter De of the electrode 14).

[0042] The influence of the presence or absence of a flow of the liquid sample on the sensor output was evaluated for the biosensor 1 of the first embodiment by conducting an experiment similar to that for the comparative biosensor 9. That is, as shown in Fig. 8, the tip of the biosensor 1 was inserted into the liquid sample in the container 20 to fill the measurement space 13 with the liquid sample, and the output of the biosensor 1 was measured while stirring the liquid sample and stopping the stirring using the stirrer 22.

[0043] FIG. 9 is an explanatory diagram showing the experimental results. Similar to the experimental results shown in FIG. 5 , the sensor output dropped sharply when the biosensor 1 was inserted into the liquid sample. Then, stirring of the liquid sample was initiated while the sensor output was dropping (i.e., approximately 10 seconds after the biosensor 1 was inserted). Despite the start of stirring while the sensor output was dropping, the sensor output continued to drop and stabilize, and the output did not fluctuate even while stirring continued. Furthermore, even after repeated stops and starts of stirring, the sensor output maintained a stable value. From these experimental results, it appears that the use of the biosensor 1 of the first embodiment makes it possible to measure the concentration of the target component a without being affected by the presence or absence of a flow in the liquid sample. This is possible because, even if there is a flow in the liquid sample, the flow is barely affected by the liquid sample in the measurement space 13, allowing the electrode 14 in the measurement space 13 to form a spherical diffusion region 14s.

[0044] Furthermore, although the measurement space 13 of the biosensor 1 of the first embodiment has an extremely simple structure as shown in Figure 1 or 7, the reason why a great effect is obtained as shown in Figure 9 is thought to be due to the existence of a hydrodynamic mechanism resulting from the minute size of the electrodes 14. This point will be explained below.

[0045] All liquids have mass and viscosity. Therefore, it is known that the flow of a liquid is affected by the inertial force due to its own mass and by viscosity. The Reynolds number Re is known as an index representing the magnitude of the influence of inertial force and viscosity. The Reynolds number Re is a dimensionless number obtained by the following calculation formula. The larger the value, the stronger the influence of inertial force, and the smaller the value, the stronger the influence of viscosity. Re = (flow velocity U) * (characteristic length L) / (liquid dynamic viscosity coefficient ν). Here, the characteristic length L is a dimension that characterizes the flow. For example, in the case of a flow in a pipe, the inner diameter of the pipe is the characteristic length L. In the case of the measurement space 13 of the biosensor 1 of the first embodiment, the diameter D or height H shown in FIG. 7 is the characteristic length L.

[0046] The biosensor 1 of the first embodiment uses a minute electrode 14, so the size of the measurement space 13 can be small. Therefore, the diameter D or height H shown in FIG. 7 can be a few millimeters. Furthermore, assuming that the flow rate of the liquid sample is only a few millimeters per second, the Reynolds number Re will be a single-digit to double-digit value. A flow with a Reynolds number Re of single to double digits is dominated almost entirely by viscosity, like the flow of honey. In contrast, if the measurement space 13 does not exist, the dimension characterizing the flow (characteristic length L) would be, for example, the dimension of the container 20, and the Reynolds number Re is expected to be about two orders of magnitude larger. If the Reynolds number Re is about two orders of magnitude larger, the influence of inertial forces will be significant.

[0047] Considering the above, the reason why the biosensor 1 of the first embodiment achieves the significant effect shown in Figure 9 is thought to be due to the following hydrodynamic mechanism. First, the Reynolds number Re is large at positions far from the protruding wall 12, so the flow of the liquid sample is greatly affected by inertial forces, and the liquid sample tends to continue flowing. However, near the protruding wall 12, the Reynolds number Re becomes sufficiently small, resulting in a viscous flow. As a result, the flow of the liquid sample rapidly decelerates as it approaches the protruding wall 12, forming an area around the protruding wall 12 where there is almost no flow. As a result, it is thought that the flow of the liquid sample within the measurement space 13 is not affected by the external flow of the liquid sample. Conversely, if the diameter of the electrode 14 is large (e.g., about several mm), the flow does not decelerate much even near the protruding wall 12, and this effect extends to the interior of the measurement space 13, resulting in significant fluctuations in the sensor output. That is, the protruding wall 12 not only defines the space around the electrode 14 as the measurement space 13 from within the larger space, but also functions as a deceleration member (deceleration wall) that forms a region outside the measurement space 13 where the flow of the liquid sample slows down (deceleration field). Therefore, as long as it is possible to define the space around the electrode 14 as the measurement space 13 and to decelerate the flow of the liquid sample, it is not necessary to have the protruding wall 12 protrude from the periphery of the electrode 14. For example, a plate-like member may be provided at a predetermined distance above the electrode 14, and the measurement space 13 may be defined below the plate-like member. Even in this case, as long as the electrode 14 and the plate-like member are positioned at an appropriate distance, the plate-like member can function as a deceleration member and can be used in place of the protruding wall 12.

[0048] Furthermore, to investigate the effect of the shape of the measurement space 13 (specifically, the combination of the diameter D and height H shown in FIG. 7 ) on the stability of the sensor output, the following experiment was conducted. First, an index called the "sensor output increase rate" was introduced to quantify the stability of the sensor output. Here, the sensor output increase rate is a value indicating the ratio of the sensor output increased by stirring the liquid sample relative to the sensor output before stirring. Then, the diameter D and height H of the measurement space 13 were changed, and the sensor output increase rate was calculated for various combinations.

[0049] FIG. 10 is an explanatory diagram showing the experimental results thus obtained. FIG. 10( a) shows a formula for calculating the increase rate R of the sensor output. The increase rate R was calculated by dividing the increase amount dJ of the sensor output due to stirring by the sensor output J before stirring and multiplying the result by 100. Furthermore, to enable a rough understanding of the influence of the shape of the measurement space 13 on the increase rate R, the increase rate R was judged into four levels. That is, as shown in FIG. 10( b), the increase rate R was judged into four levels: "A" when the increase rate R was 10% or less; "B" when the increase rate R was greater than 10% but less than 20%; "C" when the increase rate R was greater than 20% but less than 30%; and "D" when the increase rate R was greater than 30%. 10(c), the diameter D of the measurement space 13 was set to either 2 mm or 4.5 mm, and the rate of increase R of the sensor output was calculated when the height H was increased in 1 mm increments from 0 mm to 6 mm for each diameter D. The diagonal dashed line in the table of FIG. 10(c) represents the condition where the diameter D and height H of the measurement space 13 are equal.

[0050] 10(c), the experimental results show that the region below and to the left of the dashed line (i.e., under the condition that the height H is smaller than the diameter D) is rated D, and the region above and to the right of the dashed line (i.e., under the condition that the height H is larger than the diameter D) is rated A. From this, it is considered that it is desirable for the height H of the measurement space 13 to be equal to or larger than the diameter D, and more desirably, to be equal to or larger than 1.5 times the diameter D if possible.

[0051] 10(c) is for the case where the measurement space 13 is cylindrical, but if the measurement space 13 is prismatic, it is sufficient to assume a dimension equivalent to the diameter D. That is, the diameter of a circle having the same area as the area of ​​the base of the pillar shape is calculated, and this diameter is used instead of the diameter D.

[0052] As described above in detail, in the biosensor 1 of the first embodiment, the limiting permeation layer, which is essential in general biosensors, is replaced with a spherical diffusion region 14s formed around the electrode 14. This eliminates the problems caused by the limiting permeation layer (i.e., the need for calibration of the sensor output and the significant limitations on the conditions of use). In addition, because the electrode 14 is attached within the measurement space 13, even when the biosensor 1 is inserted into a flowing liquid sample, it is possible to accurately measure the concentration of the target component a in the liquid sample without being affected by the flow.

[0053] Various modifications can be considered for the biosensor 1 of the first embodiment described above. For example, as described above with reference to Figures 1 and 7, the biosensor 1 of the first embodiment has a cylindrical protruding wall 12 protruding from the insertion portion 10, and the measurement space 13 is formed inside the protruding wall 12. However, a cylindrical recess may be formed at the tip of the insertion portion 10, and the measurement space 13 may be formed inside the recess.

[0054] 11A and 11B are explanatory diagrams of a biosensor 1 according to a first modification of the first embodiment. Fig. 11A shows the external shape of the tip of the biosensor 1 according to the first modification, and Fig. 11B shows its cross-sectional shape. Of course, a recess may be formed on the side of the insertion portion 10, and the measurement space 13 may be formed within the recess. This biosensor 1 according to the first modification also uses a mechanism similar to that of the biosensor 1 according to the first embodiment described above, making it possible to realize a biosensor 1 that does not require calibration, has fewer restrictions on usage conditions, and is not affected by the flow of a liquid sample.

[0055] In the biosensor 1 of the first embodiment, one end of the measurement space 13 is open, and when the biosensor 1 is inserted into a liquid sample, the liquid sample flows into the measurement space 13 from the open end face. However, the one end of the measurement space 13 may also be closed by the ceiling portion 12r, and multiple communication openings 13o may be formed in the protruding wall 12.

[0056] FIG. 12 is an explanatory diagram of a biosensor 1 according to a second modified example of the first embodiment. FIG. 12(a) shows the external shape of the tip portion of the biosensor 1 according to the second modified example, and FIG. 12(b) shows the cross-sectional shape. In the second modified example illustrated in FIG. 12, three communication ports 13o are formed in the protruding wall 12, and the total opening area of ​​these communication ports 13o is greater than the surface area Sd of the spherical diffusion region 14s formed around the electrode 14. When the biosensor 1 according to the second modified example is inserted into a liquid sample, the liquid sample flows into the measurement space 13 through the multiple communication ports 13o, forming a spherical diffusion region 14s around the electrode 14. As a result, by using a mechanism similar to that of the biosensor 1 according to the first embodiment, it is possible to realize a biosensor 1 that does not require calibration, has fewer restrictions on usage conditions, and is not affected by the flow of the liquid sample.

[0057] Furthermore, in the biosensor 1 of the first embodiment, the measurement space 13 has been described as being formed inside the protruding wall 12. However, the measurement space 13 may be formed inside the cover member 12c by forming a dome-shaped cover member 12c using a plate-like member with numerous small holes formed therein, such as a wire netting, a mesh member, or a punched metal, and covering the electrode 14 with the cover member 12c.

[0058] FIG. 13 is an explanatory diagram of such a biosensor 1 according to a third modified example of the first embodiment. FIG. 13(a) shows the external shape of the tip portion of the biosensor 1 according to the third modified example, in which the cover member 12c is formed of wire mesh. FIG. 13(b) shows the cross-sectional shape of the tip portion of the biosensor 1 according to the third modified example, in which the cover member 12c is formed of punched metal. In either case, a measurement space 13 is formed inside the cover member 12c. When the biosensor 1 according to the third modified example is inserted into a liquid sample, the liquid sample flows into the cover member 12c through the mesh of the wire mesh, the openings in the mesh member, and the small holes in the punched metal, filling the measurement space 13 with the liquid sample, forming a spherical diffusion region 14s around the electrode 14. Note that the portions through which the liquid sample passes through the cover member 12c, i.e., the openings in the wire mesh or mesh member, and the small holes in the punched metal, each correspond to the communication ports 13o of the second modified example (see FIG. 12), and the total area of ​​the plurality of communication ports 13o is greater than the surface area Sd of the spherical diffusion region 14s formed around the electrode 14. With this type of biosensor 1 of the third modified example, it is possible to realize a biosensor 1 that does not require calibration, has fewer restrictions on usage conditions, and is not affected by the flow of the liquid sample, using a mechanism similar to that of the biosensor 1 of the first embodiment described above.

[0059] B. Second Example In the biosensor 1 of the first example and the various modified examples of the first example, the spherical diffusion region 14s formed around the electrode 14 is described as not contacting the inner surface of the member (e.g., the protruding wall 12, the ceiling portion 12r, or the cover member 12c) that forms the measurement space 13. However, the diffusion region 14s formed by the electrode 14 may be in contact with the inner surface of the member that forms the measurement space 13.

[0060] FIG. 14 is an explanatory diagram of a biosensor 1 of a second embodiment in which a diffusion region 14s formed by the electrode 14 contacts the inner surface of a member forming the measurement space 13. FIG. 14(a) shows the external shape of the tip portion of the biosensor 1 of the second embodiment, and FIG. 14(b) shows the cross-sectional shape of the tip portion. As shown in the figure, the biosensor 1 of the second embodiment has a circular disk-shaped ceiling portion 12r attached to the tip of the cylindrical insertion portion 10 with a gap of dimension Lg, and a measurement space 13 is formed between the tip of the insertion portion 10 and the ceiling portion 12r. The electrode 14 is attached to the center of the tip of the insertion portion 10. Note that in the example shown in FIG. 14, the ceiling portion 12r is supported by three thin supports 12p, but the ceiling portion 12r may be supported by other structures. Also, in Figure 14, the insertion portion 10 is cylindrical and the ceiling portion 12r is disk-shaped, but this is not limited to this, and for example, the insertion portion 10 may be a rectangular pillar and the ceiling portion 12r may be a polygonal plate-shaped member.

[0061] When the biosensor 1 of the second embodiment is inserted into a liquid sample, the liquid sample flows in through the gap between the tip of the insertion part 10 and the ceiling part 12r, filling the measurement space 13. Therefore, in the biosensor 1 of the second embodiment, the outer peripheral side surface of the measurement space 13 formed between the insertion part 10 and the ceiling part 12r serves as the communication port 13o.

[0062] Here, the gap Lg between the insertion portion 10 and the ceiling portion 12r is set to a dimension smaller than the radius of the spherical diffusion region 14s (shown by the dashed line in Figure 14(b)) that would be formed by the electrode 14 if the ceiling portion 12r were not present. Therefore, the diffusion region formed around the electrode 14 comes into contact with the ceiling portion 12r before becoming spherical and then grows in the radial direction of the ceiling portion 12r, resulting in the formation of a disk-shaped diffusion region with a thickness of Lg within the measurement space 13. The radii of the ceiling portion 12r and the insertion portion 10 are set to dimensions sufficiently larger than the radius of the disk-shaped diffusion region. Even with this biosensor 1 of the second embodiment, it is possible to achieve a biosensor 1 that is less susceptible to the influence of the flow of the liquid sample for the following reasons.

[0063] Fig. 15 is an explanatory diagram showing why the biosensor 1 of the second embodiment is less susceptible to the influence of the flow of the liquid sample. As described above with reference to Fig. 14, when the biosensor 1 of the second embodiment is inserted into a liquid sample, a disk-shaped diffusion region 14c is formed in the gap between the insertion portion 10 and the ceiling portion 12r (i.e., the measurement space 13). The circle indicated by the dashed line in Fig. 15 represents the shape of the diffusion region 14c when viewed from the tip side of the biosensor 1.

[0064] Here, the dimension Lg of the gap between the insertion section 10 and the ceiling 12r is smaller than the radius of the spherical diffusion region 14s and therefore smaller than the height H (see FIG. 7) of the measurement space 13 in the first embodiment. Therefore, the Reynolds number Re when the gap Lg between the insertion section 10 and the ceiling 12r is taken as the representative length L is smaller than the Reynolds number Re of the flow around the protruding wall 12 in the first embodiment. Therefore, the flow in the gap between the insertion section 10 and the ceiling 12r in the second embodiment is more strongly influenced by viscosity than the flow around the protruding wall 12 in the first embodiment. Therefore, even if a flow exists in the liquid sample, the flow will rapidly decelerate when it attempts to enter through the gap between the insertion section 10 and the ceiling 12r. As a result, the diffusion region 14c formed in the measurement space 13 is only slightly affected. The solid arrows in FIG. 15 conceptually represent the rapid deceleration of the flow entering through the gap between the insertion section 10 and the ceiling 12r. The dashed-dotted line in Figure 15 conceptually represents diffusion region 14d deformed by the influence of the flow entering through the gap. A comparison of diffusion region 14c and diffusion region 14d reveals that the deformation of diffusion region 14c is only slight. This level of deformation does not significantly increase the diffusion rate of target component a to electrode 14, unlike the case of Figure 6 described above. As a result, the biosensor 1 of the second embodiment can also be realized as a biosensor 1 that does not require calibration, has fewer restrictions on usage conditions, and is hardly affected by the flow of the liquid sample.

[0065] C. Third Example In the biosensor 1 of the first and second examples described above, the opening area of ​​the communication port 13o of the measurement space 13 is larger than the surface area Sd of the spherical diffusion region 14s formed by the electrode 14. However, the opening area of ​​the communication port 13o may be smaller than the surface area Sd of the spherical diffusion region 14s.

[0066] 16 is an explanatory diagram showing the cross-sectional shape of the tip portion of the biosensor 1 of the third embodiment. The biosensor 1 of the third embodiment differs from the biosensor 1 of the first embodiment described above with reference to FIG. 7 only in that the diameter D of the inner circumferential surface of the protruding wall 12 is smaller. Therefore, in the third embodiment, the area of ​​the communication opening 13o of the measurement space 13 (the end surface on the open side of the measurement space 13) is smaller than the surface area Sd of the spherical diffusion region 14s formed by the electrode 14 in the liquid sample.

[0067] In the biosensor 1 of the third embodiment, the rate at which the target component a passes through the communication port 13o due to diffusion determines the rate at which the target component a is supplied to the electrode 14. In the biosensors 1 of the first and second embodiments described above, the rate at which the target component a is supplied to the electrode 14 is determined by the diffusion region around the electrode 14, but in the biosensor 1 of the third embodiment, the rate is determined by the communication port 13o of the measurement space 13. This has the adverse effect of reducing the current value (i.e., sensor output) generated at the electrode 14, but on the other hand, it also has the effect of making the sensor output less susceptible to the influence of the presence or absence of flow. The reasons for this are as follows.

[0068] First, the concentration of the target component a in the liquid sample into which the biosensor 1 is inserted is a constant concentration (substrate concentration c). Meanwhile, in the measurement space 13, the electrode 14 converts the target component a one after another, but only a small amount of the target component a is supplied from the communication port 13o. The enzyme on the electrode 14 decomposes the target component a at a sufficiently fast rate, so the concentration at the electrode 14 is zero. Therefore, the concentration gradient in the measurement space 13 is such that the concentration at the communication port 13o is the substrate concentration c and the concentration at the electrode 14 is zero. In other words, the rate at which the target component a is supplied to the electrode 14 is the rate at which the target component a flows from the communication port 13o into the measurement space 13, and this rate is determined by the substrate concentration c of the target component a and the distance from the communication port 13o to the electrode 14 (the path length of the measurement space 13). Ultimately, the current value output by electrode 14 is determined by the substrate concentration c of the component a to be measured, the distance from communication port 13o to electrode 14, and the area of ​​communication port 13o. Therefore, with biosensor 1 of the third embodiment shown in Figure 16, it is unlikely that a situation will occur in which it is significantly affected by the flow of the liquid sample, as in the case shown in Figure 6. Furthermore, if the diameter of communication port 13o is too small, it is thought that the influence of the inner wall surface of measurement space 13 will become apparent, so it is desirable that the diameter of communication port 13o be 10 μm or more.

[0069] Of course, if there is a flow in the liquid sample at the communication port 13o, it will promote diffusion from the communication port 13o into the measurement space 13, which is thought to have some effect on the output of the electrode 14. This effect can be considered using the following model. That is, the communication port 13o forms a boundary surface 16s through which the target component a passes toward the measurement space 13, and if there is a flow in the liquid sample at the communication port 13o, the boundary surface 16s is deformed and its area increases, resulting in an increase in the output of the electrode 14.

[0070] FIG. 17 conceptually illustrates how the boundary surface of the communication port 13o deforms due to the influence of the flow in the liquid sample. The dashed line in FIG. 17 represents the undeformed boundary surface, while the dashed line in FIG. 17 represents the deformed boundary surface. Because the boundary surface is formed at the communication port 13o, the position where the outer edge of the boundary surface contacts the communication port 13o does not move regardless of whether the boundary surface is deformed or not. Therefore, even if the boundary surface is deformed due to the flow in the liquid sample, the area of ​​the boundary surface does not change significantly. Alternatively, a model can be considered in which diffusion from the communication port 13o to the measurement space 13 is promoted, resulting in an increase in the concentration of the target component a near the entrance of the measurement space 13, as if the region of substrate concentration c had slightly intruded into the measurement space 13, thereby increasing the output of the electrode 14. Even using such a model, the area of ​​the communication port 13o is small, making it unlikely that the region of substrate concentration c would extend significantly into the measurement space 13. Therefore, regardless of which model is used, the speed at which the target component a is supplied to the electrode 14 does not change significantly depending on whether there is a flow in the liquid sample, and fluctuations in the sensor output can be significantly suppressed.

[0071] On the other hand, the biosensor 1 of the third embodiment has the drawback of reducing the sensor output because the supply rate of the target component a to the electrode 14 is suppressed. However, depending on the target component a, the number of electrons n generated upon conversion at the electrode 14 may be large, or the diffusion coefficient Kd of the target component a may be large depending on the combination of the liquid sample and the target component a. Furthermore, the diffusion coefficient Kd also increases when the temperature of the liquid sample is high. In these cases, the sensor output increases, thereby compensating for the above-mentioned drawbacks. Therefore, the biosensor 1 of the third embodiment, using a mechanism similar to that of the biosensor 1 of the first embodiment described above, makes it possible to realize a biosensor 1 that does not require calibration, has fewer restrictions on usage conditions, and is not affected by the flow of the liquid sample.

[0072] Furthermore, in the above-described third embodiment, a case has been described in which the idea of ​​making the portion of the communication port 13o of the measurement space 13 determine the supply rate of the measurement target component a to the electrode 14 has been applied to the biosensor 1 of the first embodiment. However, the biosensor 1 to which the idea of ​​the third embodiment is applied is not limited to the biosensor 1 of the first embodiment, and may be applied to various modified biosensors 1 of the first embodiment and the biosensor 1 of the second embodiment. Even in these cases, it is possible to realize a biosensor 1 that does not require calibration, has fewer restrictions on usage conditions, and is not affected by the flow of the liquid sample.

[0073] Furthermore, the biosensor 1 of the various embodiments or modifications described above can also be used to measure the concentration of a component in a liquid sample flowing through a minute groove-like channel called a microfluidic device. In this case, it is desirable that the diameter of the electrode 14 is 1 mm or less (more preferably 0.1 mm or less) and the height H is 1 mm or less.

[0074] Furthermore, it is believed that the use of the biosensor 1 according to the various embodiments or modifications described above makes it possible to monitor or manage the progress of various reactions. Therefore, the biosensor 1 according to the various embodiments or modifications may be mounted in a stirring reaction vessel.

[0075] Fig. 18 is an explanatory diagram illustrating an example of a stirring reaction vessel 30 equipped with a biosensor 1. Fig. 18(a) illustrates an example of a stirring reaction vessel 30 in which the volume of the liquid to be reacted is about several liters, and Fig. 18(b) illustrates an example of a stirring reaction vessel 30 in which the volume of the liquid to be reacted is about 100 liters. Of course, a larger-scale stirring reaction vessel 30 can also be used.

[0076] As shown in FIG. 18( a) or 18(b), the stirring reaction vessel 30 includes a stirring blade 31 for promoting and homogenizing the reaction within the vessel, and a stirring motor 32 for rotating the stirring blade 31. In the case of a small-scale stirring reaction vessel 30, as shown in FIG. 18(a), a stirrer or the like can be used instead of the stirring blade 31. The stirring reaction vessel 30 may also include a heater 33 for controlling the reaction and a chemical solution inlet 34. Furthermore, in the case of a large-scale stirring reaction vessel 30, as shown in FIG. 18(b), a discharge passage 35 for discharging the reaction product may be installed. By installing one or more biosensors 1 in such a stirring reaction vessel 30, the concentration of the target component a can be stably and accurately detected even while the stirring blade 31 is stirring the liquid sample in the stirring reaction vessel 30. As a result, the progress of the reaction can be monitored accurately and in real time, enabling the reaction to be appropriately managed.

[0077] The above describes various embodiments and modifications of the biosensor 1, but the present invention is not limited to the above embodiments and modifications, and can be implemented in various forms within the scope of the gist of the present invention.

[0078] DESCRIPTION OF SYMBOLS 1...biosensor, 10...insertion portion, 11...flange, 12...projecting wall, 12c...cover member, 12p...support, 12r...ceiling portion, 13...measurement space, 13a...bottom surface, 13o...communication port, 14...electrode, 14c, 14d, 14s...diffusion region, 14z...enzyme, 15...lead wire, 15a...electric wire, 16d, 16s...boundary surface, 20...container, 21...rotor, 22...stirrer, 30...stirring reaction vessel, 31...stirring blade 31, 32...stirring motor 32, 33...heater 33, 34...feed port 34, 35...discharge passage.

Claims

1. A biosensor that measures the concentration of a predetermined component to be measured contained in a liquid sample by converting the component to a component different from the component to be measured by an enzyme immobilized on an electrode, and detecting an electrical signal generated at the electrode in association with the conversion, the biosensor comprising: an insertion part that is inserted into the liquid sample; a measurement space that is filled with the liquid sample when the insertion part is inserted into the liquid sample due to the inflow of the liquid sample; an electrode that is attached within the measurement space; and an enzyme that is immobilized on the electrode in a state in contact with the liquid sample that fills the measurement space.

2. The biosensor according to claim 1, wherein the measurement space is formed in the shape of a pillar with one end open and a bottom, and the electrode is attached to the bottom surface of the pillar.

3. A biosensor as claimed in claim 2, characterized in that the measurement space is formed by erecting a protruding wall from a position around the electrode, and surrounding the space around the electrode with the protruding wall.

4. A biosensor according to claim 2 or 3, characterized in that the height of the measurement space is equal to or greater than the diameter of a circle having the same area as the area of ​​the base of the columnar shape.

5. A stirring reaction vessel equipped with the biosensor according to any one of claims 1 to 4.

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