JP2009136526A - Analysis chip - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide an analysis sensor capable of performing noninvasive measurement while hardly imparting pain or uncomfortable feeling to a subject, effectively accelerating perspiration, making it possible to examine the presence or absence and the quantity of specific substance with a small amount of a body fluid to make it possible to measure in a short time. <P>SOLUTION: Minute prismatic recessed portions (33) are formed in the lower surface of a resinous substrate (32), and enzyme electrodes (34a, 34b) formed by immobilizing an enzyme film (40) are provided on the inner surface of each of the recessed portions (33). Sweating electrodes (37a, 37b) are provided at positions adjacent to each of the recessed portions (33), a sweating promoter (41) is provided on the lower surface of the sweating electrode (37a), and an adhesive layer (42) is provided on the lower surface of the sweating electrode (37b). <P>COPYRIGHT: (C)2009,JPO&INPIT

Description

  The present invention relates to an analysis chip. More specifically, the present invention relates to an analysis chip that has a plurality of micro-sized concave structures, collects body fluids by contacting the skin, and identifies the types and amounts of components in the body fluids. The present invention also relates to a biological material sensor using the analysis chip.

  Recent changes in human eating habits, lack of exercise, physical and mental burdens due to overwork and stress, smoking, alcohol drinking, etc. have caused damage to the immune system inherent in humans, resulting in various diseases. ing. Since lifestyles are deeply involved in the onset and progression of diseases related to this, it is generally called lifestyle-related diseases. Lifestyle-related diseases include obesity, hyperlipidemia, diabetes, hypertension, cancer, stroke, liver disease, and osteoporosis.

  In particular, the number of diabetic patients is increasing significantly worldwide. In Japan, the number of diabetic patients is said to be 6.9 million in 2005, and is said to be 170 million worldwide. Diabetes often has no subjective symptoms, so there are many people who are not treated even if they say diabetes. Without treatment, it progresses slowly in the body, causing many complications that can lead to blindness and amputation. For this reason, it is necessary to regularly check (inspect) how far the disease has progressed. You should continue to check to see if your glycemic control is good or bad and find signs of complications early.

  Diabetes mellitus is a disease that occurs when the amount of insulin secreted by the pancreas is insufficient, sugar is not used, and overflows into the blood. Therefore, it is necessary to regularly monitor blood sugar and inject an appropriate amount of insulin into the body based on the result.

  Currently, in order to monitor blood sugar, blood is actually collected and the enzyme reaction is detected electrochemically or by color. However, there are a number of concerns with blood collection. One is a physical and mental burden caused by blood invading the skin. Diabetic patients are required to measure several times a day before meals and after meals. That is, it is necessary to collect blood by inserting a needle into the skin several times a day. Blood can also be infected. In severe patients, monitoring is also required during sleep, and continuous monitoring is strongly desired. Therefore, a sensor that can monitor blood glucose level without piercing the skin (that is, non-invasively) is strongly desired.

(Invention of Patent Document 1)
An analysis sensor for measuring a blood glucose level is disclosed in Japanese Patent Application Laid-Open No. 9-5296 (Patent Document 1). As shown in FIG. 1A, this analysis sensor has a biosensor chip 12 attached in a recess provided on the inner surface of the holding member 11. As shown in FIG. 1 (b), the biosensor chip 12 is formed by forming a pair of comb-like electrodes 14a and 14b on the lower surface of the substrate 13 and covering the surface with a protective electrode 15. The enzyme membrane 16 and the separation membrane 17 are laminated on each other.

  This analytical sensor is used with the surface on which the biosensor chip 12 is provided being pressed against the surface of the skin 18. The analysis sensor pressed against the skin 18 collects the sweat secreted from the skin surface through the separation membrane 17, reacts the enzyme in the enzyme membrane 16 with the components contained in the sweat, and generates an electrical signal at that time. Is detected by comb-like electrodes 14a and 14b. Then, the type and amount of the sweat component are specified based on the detection signal. By measuring the amount of glucose in sweat noninvasively in this way, the blood glucose level can be calculated by the analytical sensor.

  However, the body surface, particularly the skin surface such as the arm, is curved, and when viewed at the micro level, the skin surface has fine irregularities. In addition, since the sweat evaporates and easily escapes, in order to collect the sweat efficiently, it is necessary to improve the adhesion between the analysis sensor and the skin surface. For that purpose, it is necessary to press the sensor strongly against the skin, which gives the user discomfort and pain. In particular, when the analytical sensor is small, the separation membrane 17 and the enzyme membrane 16 become very small, so that it is easy to give pain to the skin surface.

  On the other hand, if the chip size of the analysis sensor is increased, the user's discomfort is reduced. However, as the chip size increases, the volume inside the holding member 11 naturally increases, so that the sweat is collected and reacted. Increases the amount of sweat needed. Therefore, there arises a disadvantage that the measurement time by the analysis sensor becomes long.

(Invention of Patent Document 2)
2 and 3 are an exploded perspective view of the sampling device 21 disclosed in Japanese Patent Publication No. 10-505761 (Patent Document 2) and a perspective view showing a use state thereof. As shown in FIG. 2, the sampling device 21 includes an electrode 22, a protective layer 23 that prevents the electrode 22 from coming into contact with the skin, and a pair of transfer adhesive layers that sandwich and hold the electrode 22 and the protective layer 23. 24, a foam layer 25 affixed to the outside of each transfer adhesive layer 24, a transparent top layer 26 affixed to an adhesive on the upper surface of the upper foam layer 25, and a lower foam layer 25 The adhesive protective layer 27 is attached to the adhesive on the lower surface. In addition, circular reservoirs 28 are opened in the transfer adhesive layer 24 and the foam layer 25, respectively.

  When measuring, the adhesive protective layer 27 of the sampling device 21 is peeled off to expose the adhesive of the foam layer 25, and the pair of sampling devices 21 are attached to the body surface with the adhesive as shown in FIG. . Each sampling device 21 is connected to a controller 29 having a power source. Then, by applying a voltage with different polarity to the electrodes 22 of both sampling devices 21 and applying current to the skin, the exudate is secreted from the skin and collected in the reservoir 28.

  However, in such a sampling device, since voltage is applied to the skin, damage (pain) is given to the skin in proportion to the magnitude of the voltage, that is, the current density. Since this sampling device does not consider reducing the voltage applied to the skin, it is actually a minimally invasive sampling device even if non-invasive monitoring is intended.

  If the voltage applied to the skin, i.e., the current density, is reduced, the effect on the skin can be reduced. However, if the voltage, i.e., the current density, is reduced, the speed of extraction of exudate and body fluid is reduced, and the measurement takes time. . If the measurement takes time, the blood glucose level in blood (actual blood glucose level) and the blood glucose level in sweat can no longer be correlated, and the results measured through the collected exudate and body fluid change over time in blood glucose level. There is a problem that can not be handled.

Japanese Patent Laid-Open No. 9-5296 Japanese National Patent Publication No. 10-505761

  The present invention has been made in view of the technical problems as described above, and the object of the present invention is to make it possible to perform noninvasive measurement without causing pain and discomfort to the subject, and efficiently. An object of the present invention is to provide an analysis chip that facilitates sweating or extraction of body fluids, enables inspection of the presence or amount of a specific substance with a small amount of body fluids, and enables measurement in a short time.

  An analysis chip according to the present invention is an analysis chip for measuring a specific component in a body fluid based on a signal from a first electrode, and includes a plurality of fine samples that collect body fluid on a surface in contact with the skin. It is characterized by comprising a substrate having a recess, and the first electrode formed inside the recess and having an enzyme immobilized thereon.

  In the analysis chip according to the present invention, since the enzyme is fixed to the first electrode, a specific component in the body fluid reacts with the enzyme and the signal from the first electrode changes. Therefore, the presence / absence, amount, concentration, etc. of the specific component can be measured based on the signal from the first electrode. Moreover, since this analysis chip is provided with a plurality of fine recesses for collecting body fluid on the contact surface of the substrate with the skin, it is difficult to give pain or discomfort even when the analysis chip is pressed against the skin. In addition, since the first electrode is formed inside the minute recess, the inspection can be performed in a short time with a small amount of body fluid.

  An embodiment of the analysis chip according to the present invention is characterized in that a second electrode is disposed in the vicinity of the recess on the surface of the substrate that contacts the skin. According to such an embodiment, since the second electrode is disposed in the vicinity of each of the plurality of fine recesses, the distance between the second electrodes is shortened, so that a strong current density can be obtained with a small voltage. . In addition, when a voltage equivalent to the conventional voltage is applied to the skin using the second electrode, the current density becomes larger than the conventional voltage, so that further secretion of body fluid can be promoted and the time required for the measurement can be shortened.

  For example, if a sweat enhancer is provided on at least a part of the second electrode, the sweat is promoted more effectively by applying a voltage to the second electrode and injecting the sweat enhancer into the body. Can do.

  In another embodiment of the analysis chip according to the present invention, the second electrode extends from a surface located outside the recess to an inner surface of the recess on the surface of the substrate in contact with the skin. Formed in the region. According to such an embodiment, since the contact area of the second electrode to the skin can be increased, the electrical resistance between the second electrode is reduced, and the amount of sweat is increased by increasing the current flowing through the skin. Can do.

  In another embodiment of the analysis chip according to the present invention, the second electrode is composed of a positive electrode and a negative electrode, and the distance between the adjacent positive electrode and the negative electrode is 200 μm. It is characterized by the above. According to such an embodiment, since the distance between the positive and negative electrodes is greater than the thickness of the epidermis, the resistance in the direction along the epidermis is greater than the resistance in the thickness direction of the epidermis, and current flows to the dermis with sweat glands. Can increase the sweating promoting effect.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that the arrangement pitch of the recesses is 600 μm or more. Since the pitch of the sweat glands is about 600 μm, if the arrangement pitch of the recesses is set to 600 μm or more, it can be made substantially larger than the pitch of the sweat glands, and at least one sweat gland is located in one recess. Body fluid can be collected.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that an adhesive layer is provided on the surface of the substrate that contacts the skin. According to this embodiment, the analysis chip can be attached to the surface of the skin by the adhesive layer and can be easily fixed.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that the pressure applied to the skin during measurement is 80 mmHg or less. According to this embodiment, since the pressure for pressing the analysis chip against the skin is 80 mmHg or less, it is difficult to give pain and discomfort to the subject.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that the substrate is provided with a flow path for draining the body fluid collected in the recess to the outside of the recess. According to this embodiment, when body fluid enters the recess, the air in the recess can be discharged from the flow path to the outside. Furthermore, if the place where the body fluid spouts from the skin is blocked, the physical burden on the subject increases, but the physical burden can be reduced by providing a flow path.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that the lower end of the first electrode is located above the lower surface of the substrate. According to such an embodiment, the sweating start time is appropriately measured using the detection electrode, or the sweating time after the sweating electrode is turned on is calculated from various clinical trials, and the sweating start time is used as the sweating start time. The perspiration rate can be determined by measuring the time from when the body fluid reaches the lower end of the first electrode.

  Still another embodiment of the analysis chip according to the present invention is characterized in that the concentration of the specific component is calculated from the time from the start of the reaction obtained from the first electrode until the predetermined signal intensity is obtained and the signal intensity. It is said. The sweating rate can be known from the time from the start of the reaction obtained from the first electrode until the predetermined signal intensity is obtained, that is, the time from the start of measurement until the body fluid reaches the lower end of the first electrode, Further, since the amount of the specific component that reacts with the enzyme per unit time (reaction rate) can be known from the signal intensity, the concentration of the specific component can be determined from the perspiration rate and the reaction rate.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that the amount of the specific component is calculated from the integrated value of the signal intensity obtained from the first electrode. Since the amount of the specific component that reacts with the enzyme per unit time (reaction rate) can be known from the signal intensity obtained from the first electrode, the amount of the specific component can be obtained by integrating this amount. .

  Yet another embodiment of the analysis chip according to the present invention is characterized in that the concave portion has a prism shape or a lens shape. According to such an embodiment, the shape of the recess can be simplified. In particular, the prism-shaped recess is formed of a flat surface, so that the first electrode can be easily manufactured.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that the substrate is made of a synthetic resin. According to this embodiment, since the analysis chip can be bent, it becomes possible to fit the analysis chip tightly on a curved skin surface such as an arm.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that the substrate is formed of a thermosetting resin or an ultraviolet curable resin. According to such an embodiment, the substrate can be formed by an injection molding method, a stamper method, or the like, and the manufacturing efficiency of the analysis chip is improved.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that a battery is built in the substrate. According to the analysis chip of the present invention, since the current flowing from the second electrode to the skin can be reduced, a small battery with a small battery capacity can be used, and the battery can be built in the substrate. In addition, since the battery capacity can be reduced, the cost of the battery is reduced, and the battery can be built in the analysis chip. By these, the burden of the size or cost of the sensor or device can be reduced. In the case where a battery is built in the analysis chip, the second electrode can be automatically turned on when it comes into contact with the skin.

  Yet another embodiment of the analysis chip according to the present invention is characterized in that it comprises a fixture for wrapping around an arm or finger for fixation. According to this embodiment, the analysis chip can be attached to the skin surface having a different shape depending on the subject, and leakage when collecting body fluid can be prevented.

  A glucose sensor according to the present invention includes an analysis chip according to the present invention and an analysis apparatus main body that analyzes a signal of the analysis chip. The first electrode of the analysis chip has a specific component derived from a living body and a specific component. It is characterized by fixing an enzyme that binds automatically.

  According to the biological material sensor of the present invention, specific components derived from a living body such as glucose and cholesterol can be measured, and lifestyle-related diseases can be monitored. Or it can be used as a health check device.

  The means for solving the above-described problems in the present invention has a feature in which the above-described constituent elements are appropriately combined, and the present invention enables many variations by combining such constituent elements. .

  Hereinafter, preferred embodiments of the present invention will be described with reference to the accompanying drawings.

(Structure of the first embodiment)
Hereinafter, the structure of the analysis chip 31 according to the first embodiment of the present invention will be described with reference to FIGS. 4 is a bottom view of the analysis chip 31, and FIG. 5 is a bottom view of the state where the insulating film 36 and the sweat electrodes 37a and 37b are removed from the bottom surface of the analysis chip 31. FIG. 6 is a schematic cross-sectional view showing a cross section taken along the line XX of FIG. 4, and FIG. 7 is an enlarged cross-sectional view showing in detail the Y portion of FIG.

  The analysis chip 31 is provided with a plurality of detection units 39 and a plurality of sets of sweating electrodes 37 a and 37 b on the lower surface of the substrate 32. The substrate 32 is formed by forming a substrate material such as a thermosetting resin or an ultraviolet curable resin into a sheet shape, and has a thickness that easily curves along the skin surface. A plurality of micro-sized concave portions 33 having a prism shape (pyramid shape, truncated pyramid shape) are formed on the lower surface of the substrate 32, and the concave portions 33 are regularly arranged in a lattice shape. The recess 33 is a recess for collecting human body fluid (sweat), and a pair of minute enzyme electrodes 34a and 34b (first electrodes) are provided in the respective recesses 33 so as to face each other.

  The detection unit 39 is provided with enzyme electrodes 34a and 34b in the recess 33, and at least one surface, preferably both surfaces, of the enzyme electrodes 34a and 34b are in accordance with the test substance (specific component). An enzyme membrane 40 containing an enzyme is fixed. The enzymes contained in the enzyme film 40 may be the same for all the detection units 39, or may be of different types. If a plurality of types of enzyme films 40 are used, a plurality of types of test target substances such as glucose and cholesterol can be tested at a time. However, in order not to saturate the reaction between the enzyme and the test target substance before all the test target substances have reacted with the enzyme, a sufficiently large amount of enzyme than the amount of the test target substance is applied to the surfaces of the enzyme electrodes 34a and 34b. It is fixed.

  Each enzyme electrode 34a (positive electrode) is connected by a fine enzyme electrode wiring 35a, and each enzyme electrode 34b (negative electrode) is connected by a fine enzyme electrode wiring 35b. That is, as shown in FIG. 5, the enzyme electrode wirings 35a and 35b are respectively wired on the lower surface of the substrate 32, and the branched enzyme electrode wiring 35a enters the recess 33 and is connected to the enzyme electrode 34a. The wiring 35b enters the recess 33 and is connected to the enzyme electrode 34b, and the detection units 39 are electrically connected to each other in parallel. In order to prevent the enzyme electrode wire 35a on the positive electrode side and the enzyme electrode wire 35b on the negative electrode side from crossing each other, the enzyme electrode wire 35a and the enzyme electrode wire 35b have opposite edges of the substrate 32 as shown in FIG. Just pass through. However, when a plurality of types of enzyme membranes are used, different enzyme electrode wirings are provided for each of the enzyme electrodes 34a and 34b to which the respective types of enzyme membranes are fixed so that signals of different test substances are not mixed. 35a and 35b are used. Each detection unit 39 may be provided with a switching element (not shown). Accordingly, by scanning each detection unit 39, it is possible to sense the reaction in each detection unit 39 separately, and it is possible to sense various types of substances. Furthermore, it is possible to cancel (ignore) an error that has occurred in one detection unit 39. If one substance is sensed by a plurality of detection units 39, the number of data is reduced, but the number of data is reduced. Substance sensing becomes possible.

  As shown in FIGS. 4 and 7, an insulating film 36 is formed on the lower surface of the substrate 32 (a region excluding the concave portion 33) so as to cover the enzyme electrode wirings 35 a and 35 b. On the surface of the insulating film 36, sweat electrodes 37 a and sweat electrodes 37 b extending in the column direction are provided in the middle region between the recesses 33 and both ends in the row direction. Each sweat electrode 37 a is a sweat electrode wiring 38 a. Each sweat electrode 37b is connected by a sweat electrode wiring 38b. As shown in FIG. 4, the sweat electrode wires 38a and 38b pass through the opposite edge of the substrate 32 so that the sweat electrode wire 38a and the sweat electrode wire 38b do not intersect with each other, and the sweat electrode 37a and the sweat electrode wire 38a The sweating electrode 37b and the sweating electrode wiring 38b are both formed in a comb shape. The enzyme electrode wirings 35a and 35b and the sweating electrode wirings 38a and 38b are kept insulated from each other by the insulating film 36.

  The sweating electrodes 37a and 37b and the sweating electrode wirings 38a and 38b may be arranged in a pattern as shown in FIG. In this example, on the surface of the insulating film 36, sweat electrodes 37b extending in the column direction are provided in the middle and both ends in the row direction of the detection unit 39, and one sweat electrode that is long in the row direction is provided on the edge of the substrate 32. 37a is provided. In such an arrangement pattern, when the voltage is applied to the sweating electrodes 37a and 37b, the electric field distribution is different in each detection unit 39, so it becomes difficult to control the amount of sweating and the footback to the measurement value. However, in the process of producing the sweat electrodes 37a and 37b, the shapes of the sweat electrode 37a and the sweat electrode wiring 38a are simplified, so that the production of the electrodes and wiring is facilitated.

  In the analysis chip 31, a sweat accelerator 41 such as pilocarpine is provided on the surface of one of the sweat electrodes 37a and 37b, and an adhesive layer 42 is provided on the surface of the other electrode. Specifically, when the perspiration promoter 41 is a cationic agent, the perspiration promoter 41 is provided on the positive electrode side sweat electrode, and when the perspiration promoter 41 is an anionic agent, the perspiration promoter 41 is used as the negative electrode. On the side sweating electrode.

  When the electrode is directly applied to the skin, the contact area between the electrode and the skin fluctuates due to unevenness and curvature of the skin, resulting in variations. In order to solve this problem, a gel, which is a soft dielectric, may be applied to the electrode surface, thereby stabilizing the contact area between the electrode and the skin. Therefore, if the perspiration promoter 41 and the pressure-sensitive adhesive layer 42 are made of a gel material, the adhesion with the skin can be improved and the stability of the contact area can be maintained.

  In addition, when the electrode is directly applied to the skin, there is a risk of affecting the skin (chemical reaction) depending on the electrode material. Therefore, it is preferable to use a material that does not affect the living body or a material that does not affect the living body for the sweating electrodes 37a and 37b that come into contact with the skin. Furthermore, a material that is difficult to ionize is preferable in order to avoid an influence on the reaction. For example, silver (Ag), silver chloride (AgCl), platinum (Pt), gold (Au), stainless steel (SUS), or the like is preferably used as the electrode material.

  Note that the analysis chip 31 before use protects the enzyme electrodes 34a, 34b, the perspiration promoter 41, the adhesive layer 42, and the like from being contaminated by attaching a protective sheet to the lower surface. Alternatively, the analysis chip 31 may be sealed in the package.

(how to use)
When a protective sheet or the like is attached to the lower surface of the analysis chip 31, the protective sheet or the like is peeled off and then attached to the skin surface with an adhesive layer 42 as shown in FIG. The analysis chip 31 is used by being connected to an analysis apparatus main body 43 for data analysis, a computer or the like (hereinafter referred to as an analysis apparatus main body 43 or the like). In the analysis chip 31, a direct current voltage is applied between the sweating electrodes 37a and 37b from the analyzer main body 43 and the like, and a direct current voltage is also applied between the enzyme electrodes 34a and 34b. The current flowing between the enzyme electrodes 34a and 34b is measured by the analyzer main body 43 and the like.

  The analysis chip 31 may incorporate a battery 44 such as a button battery in the substrate 32 as shown in FIG. In this case, the battery 44 can apply a voltage between the sweat electrodes 37a and 37b and between the enzyme electrodes 34a and 34b. Since the analysis chip 31 can reduce the current flowing through the skin as will be described later, it is possible to use a small battery having a small battery capacity, and thus can be built in the substrate 32. Further, the analysis chip 31 and the analysis apparatus main body 43 may be configured integrally.

  After affixing the analysis chip 31 to the skin surface in this way, when a DC voltage of several volts is applied between the sweat electrodes 37a and 37b, the perspiration promoter 41 in contact with the skin is injected into the body (iontophoresis). , Thereby promoting fluid secretion (sweat).

  When the body fluid is thus secreted from the skin, the secreted body fluid is collected in each detection unit 39. The analyzer main body 43 analyzes the glucose concentration in the body fluid based on the value of the current flowing between the enzyme electrodes 34a and 34b, and the analysis result is displayed on the analyzer main body 43 and the like. Therefore, the analysis chip 31 and the analyzer main body 43 constitute a glucose sensor.

(Measurement principle)
FIG. 11 is a diagram for explaining a method for analyzing glucose concentration, and schematically shows the relationship between the elapsed time from the start of measurement and the amount of glucose reacted per unit time (mg / sec). Time T0 is a sweating start time, and at this time, the inside of the detection unit 39 is empty. When the measurement starts, a current flows through the sweat electrodes 37a and 37b, the sweat accelerator 41 is injected into the skin, and body fluid is secreted from the skin. The time at this time is the sweating start time TO. The body fluid secreted from the skin enters each detection unit 39. The sweating start time TO can be obtained by measuring using a detection electrode (see the fourth embodiment) or by calculating the time for sweating after the start of measurement from various clinical trials.

  Even when the body fluid enters the detection unit 39, the body fluid does not contact the enzyme electrodes 34a and 34b until the time (T1) when the body electrodes reach the lower ends of the enzyme electrodes 34a and 34b, so that glucose is not detected. Therefore, the amount of glucose reacted with the enzyme per unit time (hereinafter referred to as glucose reaction rate) is maintained at zero until time T1.

  When the body fluid rises in the detection unit 39 and reaches the lower ends of the enzyme electrodes 34a and 34b (time T1), glucose contained in the body fluid starts to react with the enzyme. At this time, the glucose contained in the body fluid is changed to gluconic acid by the action of the enzyme, and the mediator (electron carrier) is reduced by the electrons released by the glucose. The reduced mediator carries electrons to the positive electrodes of the enzyme electrodes 34a and 34b, and causes a current to flow between the enzyme electrodes 34a and 34b. Therefore, since the glucose reaction rate is proportional to the current flowing through the enzyme electrodes 34a and 34b, the glucose reaction rate can be obtained by measuring the current flowing between the enzyme electrodes 34a and 34b.

  Even if the bodily fluid in the detection unit 39 rises higher than the lower ends of the enzyme electrodes 34a and 34b, the contact area between the bodily fluid and the enzyme membrane 40 is initially small, so it does not react with the enzyme more than the amount of glucose reacted with the enzyme. The amount of glucose is higher. Therefore, as the body fluid interface rises, the reaction between glucose and the enzyme proceeds quickly, and the glucose reaction rate gradually increases. That is, the amount of reaction increases as the sample penetrates into the new enzyme layer (from time T1 to time T2). When glucose and the enzyme are in an equilibrium state, that is, when the amount of sugar diffused into the enzyme layer and the amount of enzyme reaction are stable (equilibrium), a certain amount or all of the sugar in the enzyme layer is decomposed (time) T2) Thereafter, only the glucose in the body fluid that has newly entered the detection unit 39 reacts with the enzyme, so that the glucose reaction rate becomes constant.

  Therefore, when the change of the glucose reaction rate is schematically represented, as shown in FIG. 11, the glucose reaction rate was initially maintained at zero, and then gradually increased and then maintained constant at a certain value α again. It is.

Therefore, the glucose concentration in the body fluid (the amount of glucose contained in the body fluid in a unit volume) is calculated as follows. First, the time ΔT = T1−T0 from the start of sweating until the glucose reaction rate starts to increase is the time required for the body fluid to reach the lower ends of the enzyme electrodes 34a and 34b, and the enzyme electrode 34a, Since the volume Vg below the lower end of 34b is known, the perspiration amount per unit time in one detection unit 39, that is, the perspiration rate β is
β = Vg / ΔT
It is obtained by On the other hand, the glucose reaction rate α when stabilized at a constant value represents the amount of glucose contained in the body fluid sent into one detection unit 39 during a unit time as described above. Therefore, the glucose concentration in body fluid is
Glucose concentration = α / β = (α · ΔT) / Vg
Can be obtained. Moreover, if the glucose reaction rate is integrated over time, the amount of glucose can also be determined.

  The glucose concentration and glucose amount thus determined can be useful for the prevention and diagnosis of diabetes and the like.

(Structure to increase sweating amount)
With the currently available sweating / recovery chips, it took a long time to obtain the required amount of body fluid. FIG. 12B is a diagram showing an example of the relationship between the elapsed time and the amount of collected body fluid (sweat volume) in a sweat / collection chip currently on the market. As shown in FIG. 12 (a), the perspiration promoter 41 of the perspiration electrode 37a and the adhesive layer 42 of the perspiration electrode 37b are placed on the surface of the skin 46 at a distance of about 3 cm or more, and the perspiration electrode 37a, A voltage of 15 to 30 V is applied between the DC power supply 45 between 37b and the relationship between the elapsed time (voltage application time) and the amount of collected body fluid is measured. According to FIG. 12 (b), the amount of collected body fluid per unit time is about 5 μl / min, and when this is converted into the amount of collected body fluid per unit area, about 0.8 to 1.0 μl / (min · cm ² )). In a commercially available chip, the chip volume is 50 to 100 μl including the volume that is not directly related to the reaction such as a flow path, and the area of the sweat collecting portion is reduced by pressing the skin, so it is considered to be about 5 cm ^2. It is done. Therefore, with the currently available sweating / collecting tips, it takes 10 minutes or more to fill the necessary volume with sweat.

  In order to increase the amount of perspiration and shorten the body fluid collection time, the current or current density flowing through the skin may be increased by the perspiration electrodes 37a and 37b. FIG. 13 shows the results of measurement of the amount of collected body fluid when the current flowing through the skin is 1.5 mA and the temporal change of the amount of collected body fluid when the current is 0.75 mA using a commercially available sweating / collecting chip. FIG. From this measurement result, it can be seen that the amount of sweat increases as the current flowing through the skin increases.

In order to increase the amount of sweat by increasing the current flowing through the skin, the voltage applied to the sweat electrodes 37a and 37b may be increased. However, if the applied voltage is increased or the current flowing through the skin is increased, the human body may be proportionally affected. It is said that the influence of current and voltage on the human body is surely safe if it is 20 V or less in the case of voltage.
(Http://www.mre.aist.go.jp/db-uwt1/DUCK/full/uwc2232.htm)
In the case of current density, 0.1 μA / cm ³ is said to be safe (February 2005 issue of the Institute of Electronics, Information and Communication Engineers “Summary of health effects due to electromagnetic fields”).

  If the voltage or current applied to the skin is too strong, the resistance of the skin (the epidermis) will increase and Joule heat will increase, which may cause burns.

  FIG. 14 shows a cross section of the skin pressed against the sweat electrodes 37a and 37b. The sweat electrode 37a is applied to the surface of the epidermis 47a (thickness Th1 = 0.06 to 0.2 mm), and the hair root 48 and sweat glands 49 are formed on the dermis 47b (thickness Th2 = 1 to 4 mm) below the epidermis 47a. Existing.

  FIG. 15 is an equivalent circuit model of the skin shown in FIG. In FIG. 15, Rax is the equivalent resistance of the epidermis 47a between the sweating electrodes 37a and 37b in the direction parallel to the skin surface, Ray1 and Ray2 are equivalent resistances in the thickness direction of the epidermis 47a, and Rbx is in the direction parallel to the skin surface. The equivalent resistance of the dermis 47b between the sweat electrodes 37a and 37b, and Rby1 and Rby2 are equivalent resistances in the thickness direction of the dermis 47b. The epidermis 47a (stratum corneum) has a low moisture content and therefore has a resistivity of about 10 MΩ / 100 μm, and the dermis 47b contains a lot of water and easily conducts electricity. Accordingly, the resistances Rax, Ray1, and Ray2 of the epidermis 47a are large, and the resistances Rbx, Rby1, and Rby2 of the dermis 47b are small.

  Therefore, in order to apply a voltage to the skin surface and sweat efficiently, the current applied to the dermis 47b can be passed efficiently without increasing the voltage applied to the sweat electrodes 37a and 37b, while the resistance of the epidermis 47a is reduced. A device for reducing the heat generation (Joule heat) in the skin 47a is required.

  Since the resistance Rax of the epidermis 47a varies depending on the distance P between the sweat electrodes 37a and 37b, the distance P between the sweat electrodes 37a and 37b may be reduced in order to reduce the resistance Rax.

  However, when the distance P between the sweat electrodes 37a and 37b is shortened and the resistance Rax is smaller than the resistances Ray1 and Ray2 in the thickness direction, an equivalent circuit model of the skin is as shown in FIG. Therefore, as indicated by arrows in FIG. 16, most of the current flows along the skin 47a (Rax). In particular, when the distance P between the sweat electrodes 37a and 37b is shorter than 200 μm, the resistance value becomes Rax <Ray1 and Ray2, and the current flows only through the epidermis 47a and does not easily reach the dermis 47b where the sweat glands 49 are present. It can be suppressed. Therefore, the distance P between the adjacent sweat electrodes 37a and 37b is desirably 200 μm or more.

  However, when the skin 47a is thin and wet like the lips, the resistances Rax, Ray1, and Ray2 of the skin 47a become small, and therefore, any resistance Rax, Ray1, Ray2, Rbx, Rby1, and Rby2 become small. Therefore, even if the distance P between the sweating electrodes 37a and 37b is shortened, the current flows as shown in FIG. 17, and even if the voltage is the same, the current flowing through the dermis 47b is increased, so that sweating can be promoted.

  However, the distance between the sweat electrodes 37a and 37b is limited from another aspect. In order to collect body fluid with each detection unit 39, it is desirable that at least one sweat gland 49 is opposed to all the detection units 39. Therefore, it is desirable that the arrangement pitch of the detection units 39 is larger than the pitch of the sweat glands 49. It is said that there are more than 250 sweat glands 49 per square centimeter ("Sweet common sense / insane sense" written by Tokuo Ogawa, Kodansha). That is, the pitch of the sweat glands 49 is 600 μm or more. Therefore, if the arrangement pitch of the recesses 33 is 600 μm or more, it can be expected that at least one sweat gland 49 is included in each detection unit 39, and body fluid can be collected by each detection unit 39. In the analysis chip 31, since the arrangement pitch of the detection units 39 is equal to the pitch of the sweat electrodes 37a and 37b, the distance P between the sweat electrodes 37a and 37b is preferably 600 μm or more.

  The value P = 600 μm is a value sufficient to increase the amount of sweat by reducing the distance P between the sweat electrodes 37a and 37b. That is, when the distance P is 600 μm, R = ρ × (d / S), V = RI (R: resistance value, ρ: resistivity, d: resistance length, S: resistance cross-sectional area, V: In order to obtain the same current value from the voltage (I: current value), the voltage may be about 1/50 of the conventional value.

  Therefore, if the distance P between the sweating electrodes 37a and 37b or the arrangement pitch of the detection units 39 is 600 μm or more, the sweating efficiency can be improved by miniaturizing the detection units 39 of the analysis chip 31.

  Next, the contact area between the substrate 32 and the sweat electrodes 37a and 37b and the skin will be described. FIG. 18 is a diagram showing the relationship between the area of the electrode brought into contact with the skin and the resistance (equivalent resistance of the skin). As can be seen from FIG. 18, the resistance value decreases as the electrode area increases. Therefore, it can be seen that if the electrode areas (contact areas) of the sweat electrodes 37a and 37b are increased, the resistance of the epidermis 47a is reduced, and a large current is passed through the dermis 47b to increase the sweat efficiency. In addition, since the resistance of the skin 47a is reduced, Joule heat generated in the skin 47a can be reduced, and skin damage (burn) can be prevented.

FIG. 19 shows a case where there is only one detection unit 39 (concave portion 33). In such an analysis chip, since there is a relatively large recess, if there is unevenness on the skin, a gap is easily generated between the skin and the analysis chip unless pressure is applied to the analysis chip. Body fluid 53 leaks. In order to eliminate this gap, it is necessary to apply a large pressure to the analysis chip and press it against the skin. If this happens, the subject will feel uncomfortable on the skin, causing a mental and physical burden. In addition, the period of the unevenness | corrugation of skin is about 100-340 micrometers, and the depth is 15-170 micrometers.
(Http://www.ymdlab.mce.uec.ac.jp/thesis/2006/A27mashitayukari.pdf)
Therefore, such a situation occurs when the size of the detection unit 39 is larger than about 1 mm and the depth is several hundred μm or more.

  On the other hand, in the analysis chip 31 of the first embodiment, the minute detection units 39 are arranged on the lower surface of the substrate 32 at a constant interval, so that they are in close contact with the lower surface of the substrate 32 as shown in FIG. Increases skin area. The frequency of measurement errors can be obtained without applying a large pressure to the analysis chip 31 and pressing it against the skin by performing an analysis using only the body fluid 53 collected from the detection units 39 distributed on the lower surface of the substrate 32. Can be drastically reduced. Further, when data processing is performed by the detection unit 39 that has collected body fluid, the reliability of the data is improved as the number of locations where body fluid can be collected increases. The arrangement pitch of the detection units 39 is preferably 100 μm, but since the pitch of the sweat glands 49 is about 600 μm, the arrangement pitch of the detection units 39 is preferably larger than 600 μm and 1 mm or less, and particularly about 600 μm is the best. It is believed that there is. Further, the measurement error can be reduced by averaging the data obtained by the plurality of detection units 39.

  FIGS. 21A, 21B, and 21C show the substrate 32 having a flat bottom surface and an electrode 54 formed on the entire bottom surface. When the electrode 54 formed on the entire lower surface of the substrate 32 is brought into contact with the surface of the skin 46, if the pressing force is zero, the contact area with the skin 46 is small as shown in FIG. When the substrate 32 is pressed against the skin 46, the contact area with the skin 46 is the area A of the electrode 54 as shown in FIG. Further, when the substrate 32 is strongly pressed, as shown in FIG. 21 (c), the substrate 32 is pushed into the skin 46 and also contacts the end surface. Therefore, if the contact area of the end surface is a, the total contact area is A + 2 × a.

  On the other hand, in the case of the analysis chip 31 of this embodiment, if the pressing force against the skin 46 is zero, the contact area with the skin 46 is small as shown in FIG. When the analysis chip 31 is pressed against the skin, the contact area is the area A of the sweat electrodes 37a and 37b on the lower surface of the substrate, as shown in FIG. Furthermore, when the analysis chip 31 is strongly pressed, as shown in FIG. 22C, the substrate 32 is pushed into the skin 46 and contacts the end surface, and further, the skin 46 enters the recess 33, so that the contact area of the end surface Is a and the number of the concave portions 33 is N, the total contact area is A + 2 × (N + 1) × a, and the contact area becomes large.

  Therefore, in the analysis chip 31 of this embodiment, when strongly pressed against the skin, the contact area is larger than that of the flat substrate 32, and the pressure applied to the skin is reduced to reduce the physical burden on the subject. can do.

  Further, in the analysis chip 31 of the present embodiment, the contact area between the analysis chip 31 and the skin even when compared with one with a single detection unit (see FIG. 19) or one with a flat bottom surface (see FIG. 21). Becomes larger. Accordingly, the contact area between the sweating electrodes 37a and 37b and the skin increases accordingly. When the contact area between the sweat electrodes 37a and 37b and the skin increases, the cross section through which the current flows increases, so that the resistance of the skin decreases. Therefore, the current flowing from the sweat electrodes 37a and 37b to the skin is increased, and sweating is further promoted, and heat generation due to Joule heat is reduced, and damage (burns) to the skin can be reduced.

Next, the pressing force of the sweat electrodes 37a and 37b on the skin will be described. FIG. 23 shows the result of measuring the relationship between the pressing force and the skin resistance when a circular electrode (electrode area: 12.3 cm ² ) is pressed against the skin with a bar-shaped tension gauge. According to this measurement result, although the variation is large, it can be understood that the skin resistance tends to decrease as the pressing force increases. This is considered to be because the thickness of the skin 47a is reduced by pressing the skin 47a.

  FIG. 24 shows the relationship between the skin pressing force and the contact area. In this method, a parabolic fixture is pressed against the skin using a bar-shaped tension gauge, and the area where the paint previously applied to the fixture is attached to the skin is calculated by image processing to obtain a contact area. According to this measurement result, it can be confirmed that the contact area increases as the skin pressing force increases.

  Therefore, when the analysis chip 31 is pressed against the skin, 47a is pressed by the sweat electrodes 37a and 37b, the thickness of the epidermis 47a is reduced, and the contact area between the sweat electrodes 37a and 37b and the skin is increased. Further, the thinness of the skin 47a and the reduction of the contact area between the sweat electrodes 37a and 37b and the skin reduce the resistances Ray1 and Ray2 in the thickness direction of the skin 47a. As a result, as shown in the equivalent circuit diagram of FIG. 25, even when the same voltage or power is supplied, the current flowing through the dermis 47b is increased, and the sweating efficiency is improved.

  As a method of pressing the analysis chip 31 against the skin, the analysis chip 31 may be pressed against the skin by attaching it to the arm or finger with a belt or the like like a biological material sensor (see FIG. 39) described later. Alternatively, a method of attaching the analysis chip 31 to the skin surface in an elastically curved state and pressing the analysis chip 31 against the skin with its elastic restoring force is also conceivable. However, since it is known that there is no influence on the human body when the pressure applied to the skin is 80 mmHg or less, it is preferable that the strength for pressing the analysis chip 31 against the skin is also 80 mmHg or less.

  As described above, the analysis chip 31 according to the present embodiment is devised in various ways, so that measurement can be performed non-invasively without damaging the skin or causing discomfort. In addition, sweating can be efficiently promoted, and the presence / absence and amount of a specific substance can be inspected with a small amount of body fluid, so that measurement can be performed in a short time.

(Production method)
Next, a method for manufacturing the analysis chip 31 will be described. FIG. 26 is a schematic cross-sectional view illustrating an example of a method for manufacturing the analysis chip 31. Hereinafter, the manufacturing process of the analysis chip 31 will be described with reference to FIG.

  What is shown in FIG. 26A is a stamper (mold) 52 having an inverted shape of the concave portion 33, that is, a prism-shaped convex portion 51. The stamper 52 can be produced by, for example, cutting the surface of a stamper material. Alternatively, the Si substrate or the like may be manufactured by anisotropic etching. Alternatively, the stamper 52 may be produced by electroforming from a photoresist master having recesses formed by photoresist.

  Next, as shown in FIG. 26B, the substrate 32 is manufactured using the stamper 52, and the concave portion 33 is formed in the substrate 32 by the convex portion 51. For example, the stamper 52 is set in a mold of an injection molding machine, and a thermosetting resin is injected into the mold to mold the substrate 32 having the recesses 33. Alternatively, an ultraviolet curable resin is dropped on a transparent resin flat plate, and this is pressed by the stamper 52 and spread over the entire resin flat plate. In this state, the ultraviolet curable resin is cured by irradiating ultraviolet rays. The substrate 32 is obtained by peeling.

  Thereafter, as shown in FIG. 26C, an electrode material is predetermined on the surface of the substrate 32 on the side where the recess 33 is provided by a screen printing method, an ink jet method, a sputtering method, a vacuum deposition method, a CVD method, or the like. A film is formed in a pattern, and enzyme electrodes 34 a and 34 b and enzyme electrode wires 35 a and 35 b are formed on the inner surface of the recess 33 and the lower surface of the substrate 32.

  Further, as shown in FIG. 26E, an insulating film 36 is formed on the lower surface of the substrate 32 by the microcontact method. That is, in this step, as shown in FIG. 26D, the insulating film 36 formed thin in advance is pressed against the lower surface of the substrate 32 and transferred, and the lower surface of the substrate 32 is covered with the insulating film 36.

  Next, as shown in FIG. 26 (f), an electrode material is formed in a predetermined pattern on the surface of the insulating film 36 by a screen printing method, an inkjet method, a sputtering method, a vacuum deposition method, a CVD method, etc. 37a and 37b and sweating electrode wirings 38a and 38b are produced.

  Then, as shown in FIG. 26 (g), the enzyme film 40 is fixed to the surfaces of the enzyme electrodes 34a and 34b by an ink jet method, and the sweat accelerator 41 is applied to the lower surface of the sweat electrode 37a, and the lower surface of the sweat electrode 37b. Then, the adhesive layer 42 is applied to obtain the analysis chip 31.

  According to such a manufacturing method, the analysis chip 31 with stable quality can be manufactured at low cost. In addition, if a plurality of analysis chips 31 are produced at a time using a large substrate and then separated into individual analysis chips 31, mass productivity is improved.

  In addition, as a method of manufacturing the concave portion on the substrate, without using the stamper as described above, the concave portion is formed by directly cutting the substrate with a cutting blade, the concave portion is formed by etching, or the concave portion is formed by photolithography technology. Or may be formed. Here, the method using a cutting blade is a method of cutting a recess in a substrate by applying pressure or heat to the cutting blade. The etching method is a plastic or surface processing technique that applies the corrosive action of chemicals, etc., and applies anticorrosion treatment only to the necessary parts of the material to be used, and removes unnecessary parts with a corrosive agent (etchant). This is a method for obtaining a recess having a desired shape. In particular, according to anisotropic etching, an anisotropic shape can be processed using a difference in etching rate depending on crystal orientation. In a method using photolithography technology, a mask (light-shielding material) on which a pattern is written is overlaid on a substrate coated with a photosensitive agent called a photoresist, and the photoresist is irradiated with light through the mask to expose the photoresist. In this method, the exposed portion is chemically corroded (etched) to form a desired recess. However, a method using a stamper is desirable from the viewpoint of manufacturing efficiency and cost.

  The analysis chip 31 of the present invention has various features as described above, and a plurality of minute recesses 33 for collecting body fluid (sweat) are provided on the substrate. Thus, measurement can be performed, and the physical burden on the subject can be reduced and the measurement time can be shortened. Moreover, since the several recessed part 33 is provided in the board | substrate 32 of 1 sheet, a number of parts can be reduced and cost can also be reduced.

  In addition, since the plurality of minute detection units 39 are formed on the resin substrate 32, the material cost of the analysis chip 31 can be reduced and can be manufactured at low cost by manufacturing a mold. Therefore, the analysis chip 31 can be reduced in cost, and the analysis chip 31 can be made disposable. In addition, since the resin substrate 32 is soft, the analysis chip 31 can bend following the curvature of the skin surface, and the body fluid can be prevented from leaking by being in close contact with the skin surface.

  Further, since the enzyme electrodes 34a and 34b are formed inside each detection unit 39, the inspection can be performed for each detection unit 39. Furthermore, if a plurality of types of enzymes are used, a plurality of types of test substances such as glucose and cholesterol can be tested at once, and various lifestyle-related diseases can be tested simultaneously.

  In addition, sweat electrodes 37a and 37b are provided at the portions that come into contact with the skin, a sweat accelerator is provided on one sweat electrode, and an adhesive layer is provided on the other sweat electrode, so that a voltage is applied to the sweat electrodes 37a and 37b. By applying the antiperspirant, the perspiration can be promoted and the measurement time can be shortened. Moreover, by providing the adhesive layer, the analysis chip 31 can be brought into close contact with the skin, and body fluids are less likely to leak.

  In addition, the analysis chip 31 is capable of non-invasive measurement that does not cause pain or discomfort to the subject, and also efficiently promotes sweating and performs a test for the presence or amount of a specific substance with a small amount of body fluid. It is possible to perform measurement in a short time.

  Further, in the analysis chip 31 according to the first embodiment, the detection unit 39 is configured by disposing the enzyme electrodes 34a and 34b in the micro-sized recess 33, and further, a structure unrelated to the reaction, for example, the recess 33 Since a liquid feeding part for feeding body fluid and a flow path for discharging body fluid (see FIGS. 29 to 36) are not provided, measurement can be performed with a small amount of body fluid, and the measurement time can be shortened.

  Further, in the analysis chip 31 of the first embodiment, since each detection unit 39 is connected in parallel, even when no body fluid is collected by any of the detection units 39, a test target substance such as glucose can be measured. In addition, the signal strength increases.

(Modification)
In the analysis chip 31, the shape of the recess 33 is prismatic or pyramid, but the shape of the recess 33 is not limited to the prism shape. For example, as shown in FIG. 27, the detection unit 39 may be formed by providing enzyme electrodes 34a, 34b and the like on the inner surface of a lens-shaped (hemispherical, spherical) recess 33.

  When manufacturing processes of the prism-shaped concave portion 33 and the lens-shaped concave portion 33 are compared, there are the following features. In the method of processing a recess using a cutting blade, it is necessary to control the cutting blade with high accuracy in order to obtain a spherical surface in a lens-shaped recess, but if it is a prism-shaped recess, the shape of the recess is easy to produce. .

  In the method of forming a recess by etching, in the case of a prism-shaped recess, the recess can be easily obtained by anisotropic etching of a Si substrate or the like.

  In the method using the photolithography technique, a prism-shaped concave portion or a lens-shaped concave portion can be manufactured by controlling the exposure amount to the photoresist.

  In the step of fixing the enzyme to the inner surface of the enzyme electrode or the recess, the enzyme is fixed to the curved surface in the lens-shaped recess, whereas the enzyme is fixed to the plane in the prism-shaped recess. Can be easily performed and the thickness of the enzyme membrane can be made uniform.

  Therefore, when the prism-shaped concave portion 33 and the lens-shaped concave portion 33 are compared, the prism-shaped concave portion 33 is desirable in the manufacturing process.

  FIG. 28 is a cross-sectional view showing another modification of the first embodiment. In the first embodiment, the enzyme film 40 is fixed on the surfaces of the enzyme electrodes 34a and 34b. However, in this modification, the recess 40 is filled with the enzyme 40a. According to such a modification, a large amount of enzyme 40a can be held in the small recess 33.

(Second Embodiment)
FIG. 29A is a plan view showing an analysis chip 61 according to the second embodiment of the present invention, and FIG. 29B is a sectional view thereof. In the analysis chip 61, a flow path 62 that penetrates from the top of each detection unit 39 to the upper surface of the substrate 32 is provided. In this embodiment, since the flow path 62 is provided at the upper end of the detection unit 39, when body fluid enters the detection unit 39, the air in the detection unit 39 is discharged from the flow path 62 to the outside. Further, if the place where the body fluid is ejected from the skin is blocked, the physical burden on the subject increases. However, by providing the flow path 62, such a physical burden can be reduced.

  Further, the analysis chip 61 of this embodiment does not have the sweat electrodes 37a and 37b, the sweat electrode wires 38a and 38b, and the sweat accelerator 41. Therefore, body fluids secreted by natural sweating are collected. In the case of natural sweating, there are individual differences in the amount of sweating. In the case of a subject having a large amount of perspiration, excess body fluid is discharged from the flow path 62 to the outside without overflowing in the detection unit 39.

  FIG. 30 is a schematic cross-sectional view for explaining a part of the manufacturing method of the analysis chip 61. As shown in FIG. 30A, the pin 63 protrudes from the top of the convex portion 51 provided on the stamper 52. As shown in FIGS. 30B and 30C, the stamper 52 is set on a mold 64 and a thermosetting resin is injection-molded between the stamper 52 and the mold 64, or on a flat plate. The substrate 32 is formed by pressing the ultraviolet curable resin applied to the substrate with the stamper 52 and irradiating with ultraviolet rays. In the substrate 32 thus formed, as shown in FIG. 30 (d), the concave portion 33 is formed by the convex portion 51 and the flow path 62 is formed by the pin 63. Since the subsequent steps are the same as those in FIGS. 26C to 26G showing the first embodiment, illustration and description thereof are omitted.

(Modification)
FIG. 31A is a plan view showing a modification of the second embodiment, and FIG. 31B is a sectional view thereof. In this modification, the flow path 62 provided at the upper end of the detection unit 39 does not pass through the upper surface of the substrate 32, and passes through one of the plurality of common flow paths 65 provided in parallel with the surface of the substrate 32. Opened on the side surface of the substrate 32.

  Further, as shown in FIGS. 32A and 32B, a plurality of common flow paths 65 may be gathered together at the end of the substrate 32.

  FIG. 33 is a cross-sectional view showing still another modification of the second embodiment. In this modification, a reservoir 66 (cavity) for storing extra body fluid is provided at an end portion in the substrate 32, and an end portion of the common flow path 65 is connected to the reservoir 66. According to this modification, excess body fluid overflowing from the detection unit 39 is accumulated in the reservoir 66, so that there is no possibility that the body fluid leaks from the substrate 32 and contaminates the surroundings.

  Further, as shown in FIG. 34, an air vent hole 67 may be provided in the reservoir 66. If the air vent hole 67 is provided, when the body fluid accumulates in the reservoir 66, the air in the reservoir 66 is discharged from the air vent hole 67 to the outside.

  35, enzyme electrodes 34a and 34b are provided on the inner surface of the reservoir 66, and the enzyme film 40 is fixed to the surfaces of both electrodes 34a and 34b. In this modification, the body fluid collected in the plurality of recesses 33 is sent to the reservoir 66 through the flow path 62 and the common flow path 65 by capillary action or the like, and the test substance such as glucose is detected by the enzyme electrodes 34a and 34b in the reservoir 66. Measure.

  FIG. 36 is a diagram for explaining the manufacturing method of the modified example shown in FIGS. FIGS. 36A to 36C are steps for manufacturing the substrate 32 having the recess 33 and the flow path 62, and are the same as the steps of FIGS. 30A to 30D showing the second embodiment. It is. Further, in this modification, apart from the substrate 32 of FIG. 36C, a cover member 68 having a groove-like common flow path 65 formed in the lower surface as shown in FIG. 36D is formed. The cover member 68 is preferably molded from the same resin as the substrate 32. As shown in FIG. 36 (e), the cover member 68 is overlapped and bonded to the upper surface of the substrate 32, whereby the substrate 32 in which the channel 62 and the common channel 65 are communicated as shown in FIG. 36 (f). Obtainable.

(Third embodiment)
FIG. 37 is a cross-sectional view showing the structure of an analysis chip 71 according to the third embodiment of the present invention. In this analysis chip 71, the sweat electrode 37a and the sweat electrode 37b are extended into the recess 33 in a range not contacting the enzyme electrodes 34a and 34b, respectively. When the analysis chip 71 is pressed against the skin, the skin enters the recess 33 (see FIG. 22). Therefore, if the sweat electrodes 37a and 37b are extended into the recess 33, the sweat electrodes 37a and 37b are not applied to the skin. The contact area can be increased. Therefore, the electrical resistance between the skin is reduced, and even when the voltage applied to the skin is the same, the amount of sweat can be increased by increasing the current flowing through the skin.

(Fourth embodiment)
FIG. 38 is a cross-sectional view showing the structure of the analysis chip 72 according to the fourth embodiment of the present invention. In the analysis chip 72, minute detection electrodes 73 a and 73 b for detecting body fluid are provided at the lower end of the recess 33. If a voltage is applied between the detection electrodes 73a and 73b, a current flows between the detection electrodes 73a and 73b when the body fluid reaches the detection electrodes 73a and 73b, so that the body fluid can be detected. Sweating from the time when the bodily fluid is detected (perspiration start time) until the bodily fluid reaches the lower ends of the enzyme electrodes 34a, 34b and the volume in the recess 33 between the lower ends of the detecting electrodes 73a, 73b and 34a, 34b. The speed can be obtained, and the measurement accuracy of the sweating speed can be increased.

(Fifth embodiment)
FIG. 39 (a) is a schematic diagram showing the structure of the biological material sensor 81 of the present invention. This biological material sensor 81 is provided with the analysis chip 31 detachably so as to protrude from the lower surface of the analyzer main body 43. The analyzer main body 43 has a built-in battery, and a display unit 83 for displaying the measurement result, a switch, and the like are provided on the upper surface thereof. The analyzer main body 43 is provided with a belt 82 for attachment. Here, the belt 82 is preferably made of a soft material such as a soft resin or cloth, and may be elastic or stretchable such as rubber.

  When measurement is performed, as shown in FIG. 39 (b), the analysis chip 31 is applied to the surface of the skin, the belt 82 is wrapped around the arm 84 (may be a finger), and the biological material sensor 81 is attached. To do. Since the belt 82 is made of a soft material, the biological material sensor 81 can be fitted snugly and stably on an arm or a finger of different thickness and shape depending on the subject, and body fluid leakage is prevented when collecting body fluid. be able to. The analysis chip 31 is attached to the lower surface of the analyzer main body 43 in a flat state, but is not firmly fixed to the analyzer main body 43, for example, only both ends are held. Accordingly, when the biological material sensor 81 is attached to an arm or the like, the analysis chip 31 is elastically curved along the skin surface. As a result, the analysis chip 31 is in close contact with the skin, and it is difficult for the body fluid to leak outside, and the body fluid can be reliably collected.

  Further, the analysis chip 31 can be attached to and detached from the lower surface of the analysis apparatus main body 43. Moreover, since the analysis chip 31 is mainly manufactured at a low cost using synthetic resin, it can be made disposable.

  FIG. 40 is a flowchart showing a procedure for performing measurement using the biological material sensor 81. First, the skin of the site where the biological material sensor 81 is to be mounted is cleaned by wiping or disinfecting it cleanly with alcohol or clean cotton (step S1). Note that the step of cleaning the skin can be omitted. Next, an unused analysis chip 31 is set on the lower surface of the analyzer main body 43 (step S2). Then, the biological material sensor 81 is attached to the arm 84 or the finger as shown in FIG. 39B, and the measurement is started by pressing the start button of the analyzer main body 43 (step S3).

  When the start button is pressed, the analyzer main body 43 applies a voltage to the sweat electrodes 37a and 37b of the analysis chip 31 and injects the sweat accelerator 41 subcutaneously (step S4). When the perspiration promoter 41 is injected, perspiration from the skin occurs (step S5), so that body fluid is collected in the detection unit 39 and reacted with the enzyme film 40 (step S6). And the electric current which flows into the enzyme electrodes 34a and 34b is measured, this is analyzed, the biomolecule density | concentration etc. which are contained in a bodily fluid are calculated | required (step S7), and an analysis result is displayed on the display part 83 (step S8).

  When the subject or the user of the biological material sensor 81 has confirmed the display, the analyzer main body 43 is turned off, the biological material sensor 81 is removed (step S9), and the used analysis chip 31 is discarded as necessary. .

  The biological material sensor 81 can be used as a glucose sensor for diabetes diagnosis as long as it can measure the glucose concentration and the glucose amount. In addition, the biological material sensor 81 can be used for measurement of cholesterol, hormones, drugs suspected of abuse, prohibited drugs, and the like. In particular, the biological material sensor 81 is suitable for easy measurement not only in hospitals but also at home, for prevention and diagnosis of lifestyle-related diseases such as diabetes and obesity, and health check.

FIG. 1A is a schematic cross-sectional view of the analysis sensor disclosed in Patent Document 1, and FIG. 1B is a bottom view showing the structure of the biosensor chip. FIG. 2 is an exploded perspective view of the sampling device disclosed in Patent Document 2. As shown in FIG. FIG. 3 is a perspective view showing a usage state of the sampling device disclosed in Patent Document 2. As shown in FIG. FIG. 4 is a bottom view showing the structure of the analysis chip according to the first embodiment of the present invention. FIG. 5 is a bottom view showing a state in which the insulating film and the sweat electrode are removed from the bottom surface of the analysis chip according to the first embodiment. FIG. 6 is a schematic sectional view taken along line XX in FIG. FIG. 7 is an enlarged cross-sectional view showing the Y portion of FIG. 6 in detail. FIG. 8 is a diagram showing different arrangements of the sweating electrode and the sweating electrode wiring. FIG. 9 is a perspective view showing a state in which the analysis chip according to the first embodiment is attached to the arm. FIG. 10 is a cross-sectional view showing a substrate incorporating a battery. FIG. 11 is a diagram for explaining a method of analyzing the glucose concentration in body fluid using the first analysis chip. FIG. 12A is a diagram for explaining the function of the sweating electrode, and FIG. 12B is a diagram showing the relationship between the elapsed time and the amount of collected body fluid (sweat amount) in a sweat / collection chip currently on the market. FIG. 13 is a diagram showing a temporal change in the amount of collected body fluid when the current flowing through the skin is 1.5 mA and the amount of collected body fluid when the current is 0.75 mA. FIG. 14 is a view showing a cross section of the skin pressed against the sweat electrode. FIG. 15 is a diagram showing an equivalent circuit model of the skin shown in FIG. FIG. 16 is a diagram showing an equivalent circuit model of the skin when the distance between the electrodes is short. FIG. 17 is a diagram showing an equivalent circuit model of the skin. FIG. 18 is a diagram showing the relationship between the area of the electrode brought into contact with the skin and the resistance (equivalent resistance of the skin). FIG. 19 is a schematic view showing a state in which the detection unit presses one chip against the skin. FIG. 20 is a schematic diagram illustrating a state in which the analysis chip according to the first embodiment is pressed against the skin. 21 (a), 21 (b), and 21 (c) are diagrams illustrating a state in which a substrate having a flat bottom surface is pressed against the skin with different pressing forces. FIGS. 22A, 22B, and 22C are diagrams illustrating a state in which the analysis chip according to the first embodiment is pressed against the skin with different pressing forces. FIG. 23 is a graph showing the results of measuring the relationship between pressing force and skin resistance when a circular electrode is pressed against the skin with a bar-shaped tension gauge. FIG. 24 is a diagram showing the relationship between the pressing force of the skin and the contact area. FIG. 25 is a diagram showing an equivalent circuit model of the skin. FIG. 26A to FIG. 26G are schematic cross-sectional views illustrating an example of a method for manufacturing an analysis chip according to the first embodiment. FIG. 27 is a diagram illustrating a cross section of a modified example of the first embodiment and the shape of a lens-shaped recess. FIG. 28 is a cross-sectional view showing another modification of the first embodiment. FIG. 29A is a plan view showing an analysis chip according to the second embodiment of the present invention, and FIG. 29B is a cross-sectional view thereof. FIG. 30A to FIG. 5D are schematic cross-sectional views illustrating a part of the analytical chip manufacturing method of the second embodiment. FIG. 31A is a plan view showing a modification of the second embodiment, and FIG. 31B is a sectional view thereof. FIG. 32A is a plan view showing another modification of the second embodiment, and FIG. 32B is a sectional view thereof. FIG. 33 is a cross-sectional view showing still another modification of the second embodiment. FIG. 34 is a cross-sectional view showing still another modification of the second embodiment. FIG. 35 is a cross-sectional view showing still another modification of the second embodiment. 36 (a) to 36 (f) are diagrams for explaining a manufacturing method of the modification shown in FIGS. 31 (a) and 31 (b). FIG. 37 is a sectional view showing the structure of an analysis chip according to the third embodiment of the present invention. FIG. 38 is a cross-sectional view showing the structure of an analysis chip according to the fourth embodiment of the present invention. FIG. 39 (a) is a schematic view showing the structure of the biological material sensor of the present invention, and FIG. 39 (b) is a schematic view showing a state where the biological material sensor is attached to the arm. FIG. 40 is a flowchart showing a procedure for performing measurement using a biological material sensor.

Explanation of symbols

31 Analytical chip 32 Substrate 33 Recess 34a, 34b Enzyme electrode 35a, 35b Enzyme electrode wiring 36 Insulating film 37a, 37b Sweating electrode 38a, 38b Sweating electrode wiring 39 Detection unit 40 Enzyme film 41 Sweating promoter 42 Adhesive layer 43 Analyzing device main body 44 battery 61 analysis chip 62 flow path 65 common flow path 66 reservoir 67 air vent hole 71 analysis chip 72 analysis chips 73a and 73b detection electrode 81 biological material sensor

Claims (18)

An analysis chip for measuring a specific component in a body fluid based on a signal from a first electrode,
A substrate having a plurality of fine recesses for collecting body fluid on the surface in contact with the skin;
The first electrode formed inside each of the recesses and having the enzyme immobilized thereon;
An analysis chip comprising:   2. The analysis chip according to claim 1, wherein a second electrode is disposed in the vicinity of the concave portion on the surface of the substrate that contacts the skin.   The analysis chip according to claim 2, wherein a sweating promoting agent is provided on at least a part of the second electrode.   3. The analysis according to claim 2, wherein the second electrode is formed in a region extending from a surface located outside the concave portion to an inner surface of the concave portion on a surface of the substrate in contact with the skin. Chip.   The second electrode is composed of a positive electrode and a negative electrode, and a distance between an adjacent positive electrode and negative electrode is 200 μm or more. Analysis chip.   The analysis chip according to claim 1, wherein an arrangement pitch of the recesses is 600 μm or more.   The analysis chip according to claim 5, wherein an adhesive layer is provided on a surface of the substrate that contacts the skin.   The analysis chip according to claim 1, wherein the pressure applied to the skin during measurement is 80 mmHg or less.   The analysis chip according to claim 1, wherein a flow path for draining the body fluid collected in the recess to the outside of the recess is provided in the substrate.   The analysis chip according to claim 1, wherein a lower end of the first electrode is located above a lower surface of the substrate.   The analysis chip according to claim 10, wherein the concentration of the specific component is calculated from the time from the start of the reaction obtained from the first electrode until the predetermined signal intensity is obtained and the signal intensity.   The analysis chip according to claim 1, wherein the amount of the specific component is calculated from an integrated value of the signal intensity obtained from the first electrode.   The analysis chip according to claim 1, wherein the recess has a prism shape or a lens shape.   The analysis chip according to claim 1, wherein the substrate is made of a synthetic resin.   The analysis chip according to claim 14, wherein the substrate is formed of a thermosetting resin or an ultraviolet curable resin.   The analysis chip according to claim 2, wherein a battery is built in the substrate.   The analysis chip according to claim 1, further comprising: a fixing tool that is wound around and fixed to an arm or a finger. The analysis chip according to any one of claims 1 to 17, and an analysis apparatus main body that analyzes a signal of the analysis chip,
A biological substance sensor, wherein an enzyme that specifically binds to a specific component derived from a living body is fixed to the first electrode of the analysis chip.

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