Measuring instrument provided with solid component concentrating means

The present invention relates to a measuring instrument (1) comprising a channel (60) for moving a sample liquid (BL) containing a solid component (B1) and providing a liquid reaction field and first and second electrodes (31, 32) which are used to apply voltage to the liquid reaction field. The first electrode (31) has an electron transfer interface (31a) for transferring electrons between it and the liquid reaction field when voltage is applied to the liquid reaction field via the first and second electrodes (31, 32). The measuring instrument (1) comprises concentrating means (51) for increasing the concentration of solid components at portions thereof which contact the electron transfer interface (31a) in the liquid reaction field. The concentrating means (51) preferably comprises a water-absorbing layer containing an absorbent polymer material.

TECHNICAL FIELD

The present invention relates to a measuring instrument to be used for measuring the concentration of a specific component (such as glucose or cholesterol) in a sample liquid such as blood.

BACKGROUND ART

Some methods of measuring the concentration of a specific component in a sample liquid employ electrochemical techniques for example. In these methods, a reaction system is formed from a sample liquid, an oxidoreductase and an electron transporter for example, voltage is applied to this reaction system using electrodes, and the concentration of the specific component is computed based on the resulting response current. Such a reaction system is formed for example on a biosensor provided with a reagent part comprising an oxidoreductase and an electron transporter. Because an oxidation-reduction reaction occurs between the specific component and the electron transporter in the reaction system due to the catalytic effect of the oxidoreductase, the amount of the electron transporter which becomes reduced (or oxidized) reflects the concentration of the specific component. Response current is obtained in correlation with the amount of electron movement occurring in the reaction system. Consequently, the accuracy of concentration measurement is greatly affected by the accuracy of measurement of response current.

In such methods, when blood (blood containing blood cells) is used as a sample liquid, electron movement between the electrode and the electron transporter is impeded by blood cells present on the surface of the electrode. As a result, the measured response current is reduced as the number of blood cells increases, producing measurement errors. Moreover, if the proportion of blood cells in the blood (hematocrit value) is different the measured response current will be different even if the glucose concentration is the same.

To solve these problems, methods have been proposed of separating blood cells from blood in the measuring instrument. Methods of separating blood cells include for example a method of providing a separation membrane in that part of the measuring instrument where blood or other sample liquid is introduced (see for example JP-A 8-114539 and JP-A 2002-508698), and a method of covering the surface of the electrode with a polymer membrane (see for example JP-A 6-130023, JP-A 9-243591 and JP-A 2000-338076).

However, because in methods which filter blood cells in the measuring instrument the plasma component must pass through the separation membrane it takes a long time for the plasma to reach the surface of the electrode, prolonging the measurement time. This problem can be solved by securing a large quantity of whole blood to be used, but this places a heavy burden of blood collection on the user.

DISCLOSURE OF THE INVENTION

It is an object of the present invention to control the effect of the solid component in the sample liquid and precisely measure concentration with a small amount of sample liquid while keeping the measurement time short.

In the present invention a measuring instrument is provided which is an analytical instrument comprising a channel for moving a sample liquid containing a solid component and for providing a liquid reaction field and a first and second electrode which are used to apply voltage to the liquid reaction field, wherein the first electrode has an electron transfer interface for providing electrons to the liquid reaction field or receiving electrons from the reaction field when voltage is applied to the liquid reaction field via the first and second electrodes, and which is provided with concentration means for increasing the concentration of solid components in that part which contacts the electron transfer interface in the liquid reaction field.

The concentration means comprises for example a water-absorbing layer containing an absorbent polymer material. The absorbent polymer material must be capable of absorbing enough liquid component to achieve the object of the present invention, and the amount of water absorbed must not be such as to affect the measurement results. Consequently, it is desirable to use an absorbent polymer material having a water absorption power of 10 to 500 g/g.

The measuring instrument includes for example a substrate on which are formed the first and second electrodes with a cover laid over the substrate.

The water-absorbing layer is formed as a film on at least a part of the cover which faces the electron transfer interface for example. In this case, the water-absorbing layer is preferably formed so that its dimension in the direction of thickness of the substrate without water absorption and with water absorption is 1/30 to 1/10 and ⅕ to ⅗, respectively, of the dimension in the direction of thickness of the channel. The water-absorbing layer can also be made water-soluble.

The water-absorbing layer may also be formed across all or most of the length of the channel. Such a water-absorbing layer can be the cover if the cover contains an absorbent material.

The water-absorbing layer may comprise powder of an absorbent polymer material supported on the cover. The weight average grain size of the powder without water absorption is 100 to 1000 μm for example. This is because if the average grain size in unsuitably small it is difficult to form the water-absorbing layer so that it cannot absorb enough water to achieve the object, while if the average grain size is unsuitably large it will impede the flow of water in the channel more than necessary.

The water-absorbing layer can also be placed downstream in the flow of sample liquid from the electron transfer interface in the channel. This water-absorbing layer is placed for example on either the substrate or the cover. In this case the dimension of the water-absorbing layer in the direction of flow of the sample liquid is preferably ¼ to ½ of the distance from the channel inlet to the furthest downstream point of the electron transfer interface in the direction of flow of the sample liquid so that the solid component can be concentrated as intended. For the same reasons it is desirable that during water absorption the thickness of the part having the formed water-absorbing layer be 0 to 15 μm.

The water-absorbing layer can also be formed so as to have a part formed in at least one of a location upstream from and adjacent to the electron transfer interface and a location downstream from and adjacent to the electron transfer interface. In this case, the water-absorbing layer is preferably formed so as to have both a part formed in a location upstream from and adjacent to the electron transfer interface and a part downstream from and adjacent to the electron transfer interface, and it is formed for example so as to surround the electron transfer interface.

The concentrating means can also be provided downstream in the channel from the electron transfer interface in the direction of movement of the sample liquid and can be formed by a absorption-resistant dam to impede the movement of the solid components.

The dam is formed so that the thickness of the channel where the dam part is formed is 5 to 15 μm for example so as to concentrate the solid components as intended.

The same liquid which is measured in the measuring instrument of the present instrument may classically be blood containing blood cells as the solid component. The measuring instrument of the present invention can be used for a wide range of sample liquids containing solid components, and the sample liquid which is measured is not limited to blood.

BEST MODE FOR CARRYING OUT THE INVENTION

Embodiments 1 through 5 of the present invention are explained in detail below with reference to the drawings.

First, the first embodiment of the present invention is explained using blood sugar measurement as an example with reference toFIGS. 1 through 5.

The biosensor1shown inFIGS. 1 and 2is used to measure glucose concentrations in blood, and is used mounted on concentration measuring system2(seeFIG. 5). This biosensor1includes cover5laid over rectangular substrate3with a pair of spacers40and41in between, and as shown inFIG. 3, capillary6is formed by these elements3,40,41and5.

Capillary6moves blood by means of capillary action, and has internal channel60for holding the blood. This internal channel60extends in the short direction of substrate3, and communicates with the outside via end openings61and62. End opening61is used to introduce blood into internal channel60, and when blood is moved through internal channel60end openings62is used to exhaust gas from internal channel60.

As shown inFIGS. 1 through 3, the pair of spacers40and41serve to join cover5to substrate3and to determine the dimensions of internal channel60of capillary6. The pair of spacers40and41extend in the short direction of substrate3, and are arranged with a gap between them in the long direction of substrate3.

Working electrode31and counter electrode32are formed extending in the long direction of substrate3on upper surface30of substrate3. Reagent site33is also provided on upper surface30of substrate3so as to successively transect both working electrode31and counter electrode32. The part of working electrode31which contacts reagent site33comprises electron transfer interface31a.

Reagent site33is formed as a solid which comprises an oxidoreductase and an electron transporter. Glucose oxidase or glucose dehydrogenase for example can be used as the oxidoreductase. The electron transporter is oxidized or reduced by application of voltage and reactions, and in blood sugar measurement potassium ferricyanide for example is used as the electron transporter. In this embodiment, the electron transporter is included in its oxidized form before the blood is supplied.

As shown inFIGS. 2 and 3, water-absorbing layer51is formed on one side50of cover5. This water-absorbing layer51is formed on one side50of cover5so as to face electron transfer interface31a, which is located in internal channel60on working electrode31. This water-absorbing layer51can be formed by affixing a water-absorbing sheet comprising an absorbent polymer material to cover5. This water-absorbing layer51is formed so that its thickness is for example 1/30 to 1/10 of the height (H) of internal channel60without water absorption, and so that its thickness with water absorption is ⅕ to ⅗ the height (H) of internal channel60.

A material with water absorption power of 10 to 500 g/g for example is used as the absorbent polymer material. Specific examples of the absorbent polymer material include crosslinked acrylate polymers, crosslinked vinyl alcohol-acrylate copolymers, crosslinked maleic anhydride grafted polyvinyl alcohol, crosslinked isobutylene-maleic anhydride copolymer, carboxymethylcellulose crosslinked with alkali salts, crosslinked partially neutralized polyacrylic acid and the like. Water absorbing layer51may be entirely formed of an absorbent polymer material, or it may be formed as a layer which is a mixture of an absorbent polymer material and a non-absorbent polymer material. Water-absorbing layer51can be formed by first applying a solution of an absorbent polymer material dissolved in a solvent to cover5and then drying it.

As shown inFIG. 5, concentration measuring system2includes first and second terminals20aand20b, voltage applying part21, current measuring part22, detection part23, control part24, computation part25and display part26.

First and second terminals20aand20bare provided for contact with terminals31band32bof working electrode31and counter electrode32in biosensor1when biosensor1is mounted on concentration measuring system2.

Voltage applying part21is for applying a voltage between working electrode31and counter electrode32of biosensor1via first and second terminals20aand20b. Voltage applying part21is electrically connected to first and second terminals20aand20b. Voltage applying part21includes for example a direct current power source such as a dry battery or charger.

Current measuring part22measures the current when a voltage is applied by means of voltage applying part21between terminals31band32bof working electrode31and counter electrode32.

Detection part23detects based on the current measured by current measuring part22whether or not a sample liquid has been supplied to reagent part33(seeFIGS. 1 through 3) after biosensor1has been mounted on concentration measuring system2.

Control part24controls voltage applying part21and selects whether a voltage is applied or not between working electrode31and counter electrode32.

Computation part25computes the blood sugar concentration according to the current measured by current measuring part22. Computation part25is designed so as to compute the blood sugar concentration by amperometric means for example.

Detection part23, control part24and computation part25may include multiple memories (such as ROM or RAM) connected to one CPU.

Display part26displays the computation results of computation part25along with error messages and operating procedures for example, and may include for example a liquid crystal display device.

Next, the procedures for measuring the blood sugar concentration using biosensor1and concentration measuring system2are explained.

As clearly shown inFIG. 5, biosensor1is first set on concentration measuring system2. Terminals31band32bof working electrode31and counter electrode32of biosensor1are thus connected to first and second terminals20aand20bof concentration measuring system2. This allows voltage to be applied between working electrode31and counter electrode32via first and second terminals20aand20b. Under actual measurement conditions, a constant voltage is applied between working electrode31and counter electrode32as soon as biosensor1is mounted on concentration measuring system2. The constant voltage applied between working electrode31and counter electrode32is set for example in the range of 100 to 1000 mV. In this embodiment, application of the constant voltage between working electrode31and counter electrode32is performed continuously until the response current for computing the blood sugar concentration has been measured.

Next, blood is supplied via end opening61of biosensor1. As shown inFIGS. 4A and 4B, blood BL travels by capillary action through internal channel60from end opening61to end opening62of capillary6. As clearly shown inFIG. 4B, blood BL is introduced until blood BL reaches end opening62and internal channel60of capillary6is full of blood BL. In this process reagent site33(seeFIG. 4A) is dissolved by blood BL and a liquid reaction system is formed in internal channel60. At this time water-absorbing layer51absorbs the plasma component of blood BL, and water-absorbing layer51grows in thickness. In this way, the movement of blood cells B1is impeded by water-absorbing layer51, and the concentration of blood cells B1rises on and around the surface of electron transfer interface31aof working electrode31.

Within the liquid reaction system, the glucose in blood BL is oxidized by oxidoreductase while the electron transporter becomes reduced. When a voltage is applied, the reduced electron transporter moves to the surface of electron transfer interface31aof working electrode31, supplies electrons to electron transfer interface31aand reverts to an oxidized electron transporter. The amount of electrons supplied to electron transfer interface31ais measured as a response current by current measuring part22via first and second terminals20aand20b.

Meanwhile, the response current measured by current measuring part22is monitored by detection part23, and once the response current exceeds a threshold detection part23detects that the blood has been supplied to reagent part33and reagent part33has dissolved. When detection part23has detected that the blood has been supplied, detection part23then evaluates whether a fixed time has passed since this detection.

When detection part23has determined that a fixed time has passed, a response current is measured at current measuring part22and the blood sugar concentration is computed by computing part25based on the response current. Computation of the blood sugar concentration is performed by first converting the response current to a voltage and then comparing the voltage to a previously-prepared reference curve showing the relationship between voltage values and blood sugar concentrations. The computation result of computation part25is displayed on display part26for example.

In this embodiment, when blood BL is supplied to internal channel60of capillary6the plasma component of blood BL is absorbed by water-absorbing layer51, increasing the concentration of blood cells B1on and around the surface of electron transfer interface31aof working electrode31. In this way, the area on and around the surface of electron transfer interface31ais artificially in the same state as if high-hematocrit blood BL were being supplied. Moreover, if an absorbent polymer material with a water absorption power of 10 to 500 g/g is used, water-absorbing layer51will absorb more plasma the lower the hematocrit value of blood BL. As a result, a similar high hematocrit state can be achieved around water-absorbing layer51regardless of whether the hematocrit value is high or low.

Biosensor1can also resolve issues involved in separating blood cells from blood in the measuring instrument. That is, in methods of separating blood cells the plasma component has to be passed through a separation membrane, prolonging the measurement time and reducing the amount of blood that can be used in the reaction relative to the amount supplied. In contrast, because in biosensor1there is no separation membrane or other obstruction to the progress of blood BL through internal channel60of capillary6, measurement time is not prolonged as it is with a separation membrane. Moreover, because in biosensor1most of the blood BL supplied to internal channel60of capillary6can be reacted with oxidoreductase of reagent part33, concentrations can be measured properly using only a tiny amount of blood BL.

Next, a biosensor according to the second embodiment of the present invention is explained with reference toFIG. 6.

In the biosensor1A shown inFIG. 6, water-absorbing layer51A is formed across the entire length of capillary6. This water-absorbing layer51A can be formed for example by affixing a water-absorbing sheet of an absorbent polymer material to the cover. Water-absorbing layer51A can also be formed by coating the cover with a solution of an absorbent polymer material dissolved together with an adhesive component in a solvent, and then drying the coat.

An absorbent layer can also be formed across the entire length of capillary6by making all of cover5water absorbent and having all of cover5comprise the water-absorbing layer. Such a water-absorbing layer (cover) can be formed for example by first mixing an absorbent polymer material with a resin material to make a molding material which is then subjected to resin molding.

Next, a biosensor according to the third embodiment of the present invention is explained with reference toFIG. 7.

The biosensor1B shown inFIG. 7includes water-absorbing layer51B having absorbent polymer grains. This water-absorbing layer51B, in which absorbent polymer grains51Bbare held on one side of double-sided tape51Ba, is affixed to the cover using the adhesiveness of the other side of the double-sided tape. An absorbent polymer with a weight-average grain size of 100 to 1000 μm may be preferably used.

Next, a biosensor according to the fourth embodiment of the present invention is explained with reference toFIGS. 8A and 8B.

In the biosensor1C shown inFIG. 8A, water-absorbing layer51C is formed on substrate3downstream in the direction of blood flow from electron transfer interface31aof working electrode31. However, water-absorbing layer51C may also be formed on cover5.

In this biosensor1C, as shown inFIG. 8B, when blood BL is introduced into capillary6water-absorbing layer51C expands, decreasing the spatial cross-sectional area of the part of capillary6having formed water-absorbing layer51C. As a result, the movement of blood cells B1is impeded by water-absorbing layer51C, blood cells B1accumulate near electron transfer interface31aand the concentration of blood cells B1increases around electron transfer interface31a.

In order for this to be effective, water-absorbing layer51C is preferably formed so that the distance L between water-absorbing layer51C and the upper surface of the capillary is 0 to 15 μm when capillary6is filled with blood. Moreover, in order to more reliably increase the concentration of blood cells B1around electron transfer interface31ait is desirable that dimension W1of water-absorbing layer51C in the direction of flow of blood BL be made relatively large. Dimension W1in this case is preferably set to about ¼ to ½ of the distance W2between inlet68of capillary6and downstream end31a′ of electron transfer interface31a.

A function similar to that of the water-absorbing layer51C shown inFIGS. 8A and 8Bcan be achieved with a non-(or low-) water-soluble dam part. That is, rather than causing blood cells to accumulate due to absorption of the plasma component and expansion, the cross-sectional dimensions downstream from electron transfer interface31ain capillary6can be made small by the formation of a dam before blood is supplied. This dam is preferably formed so that the distance (corresponding to L inFIG. 8B) between the dam and the substrate (or cover) is 5 to 15 μm. The dam can be formed on either the substrate or the cover.

Next, a biosensor according to the fifth embodiment of the present invention is explained with reference toFIGS. 9,10A and10B.

In the biosensor1D shown inFIG. 9, water-absorbing layer51D is formed so as to have a part adjoining electron transfer interface31aof working electrode31. As shown inFIG. 10A, water-absorbing layer51D may be arranged in two sites both upstream and downstream from electron transfer interface31a(seeFIG. 9), or as shown inFIG. 10Bit can be formed as a rectangular frame. In the configuration shown inFIG. 10A, one of the two water-absorbing layers51D may also be omitted.

The measurement of blood glucose concentration was taken as an example in the above explanation, but the present invention is applicable to measurement of other components in blood such as cholesterol, lactic acid, bilirubin and the like, as well as to sample liquids other than blood.

EXAMPLES

It is shown below that the effect of blood cells in blood on measurement of response current is reduced by the biosensor of the present invention.

In this example, a biosensor was formed with the same structure as inFIGS. 1 through 4. In this biosensor, the length L, width W and height H of internal channel60of capillary6were given as 3 mm, 1 mm and 40 μm, respectively (seeFIGS. 1 and 3). Working electrode31and counter electrode32were formed by screen printing using carbon ink (Japan Acheson “Electrodag 423SS”). Reagent site33was given a two-layer structure consisting of an electron transport layer and an enzyme-containing layer. The electron transport layer was formed by first applying 0.4 μL of a first material liquid comprising an electron transporter to substrate3, and then blow drying (30° C., 10% Rh) the first material liquid. The enzyme-containing layer was formed by first applying 0.3 μL of a second material liquid containing oxidoreductase to the electron transport layer, and then blow drying (30° C., 10% Rh) the second material liquid.

The first material liquid was prepared by mixing the liquid components in the following table 1 in numerical order from 1 to 4, letting the mixture stay for 1 to 3 days and then adding an electron transporter thereto. The second material liquid was prepared by dissolving an oxidoreductase in 0.1% CHAPS.

[Ru(NH3)6]Cl3(Dojin Chemical Laboratory “LM722”) was used as the electron transporter, and PGGGDH (800 U/mg glucose dehydrogenation activity) was used as the oxidoreductase.

In Table 1, SWN is an abbreviation for Lucentite SWN, CHAPS is an abbreviation for 3-[(3-cholamidopropyl)dimethylammonio]propanesulfonic acid, and ACES is an abbreviation for N-(2-acetamido)-2-aminoethanesulfonic acid. Coop Chemical “3150” was used as the SWN, Dojin Chemical Laboratory “KC062” as the CHAPS and Dojin chemical Laboratory “ED067” as the ACES. The ACES solution was adjusted to a pH of 7.5.

Water-absorbing layer51was formed to a thickness of 2 μm by first applying 0.1 μL of a coating material comprising an absorbent polymer to the target site of cover5and then blow drying it (30° C., 10% Rh). 7 parts by weight of absorbent polymer (Sumitomo “Aquacork”) dissolved in 100 parts by weight of methanol was used as the coating material.

Using the biosensor described above, response current was measured as a time course for three types of blood (Hct 20%, Hct 42% and Hct 69%) with different hematocrit (Hct) values and a glucose concentration of 447 mg/dL. Response current was measured 5 times for each Hct type of blood. The amount of blood supplied to internal channel60of biosensor1here was 0.5 μL, and the voltage applied between working electrode31and counter electrode32was 500 mV. The results are shown inFIG. 11.

The effect of Hct was investigated based on response current 5 seconds after the supply of blood. The results are shown inFIG. 13. InFIG. 13Hct (%) is shown on the horizontal axis and percent bias from response current at Hct 42% on the vertical axis. InFIG. 13, percent bias is shown as the average of 5 measurements.

Comparative Example 1

In this comparative example a biosensor was used having the absorbent layer omitted from the biosensor of Example 1, and a time course of response current was measured for three types of blood of differing Hct as in Example 1. Response current was measured 5 times for each Hct type of blood. The results are shown inFIG. 12. Also as in Example 1, the effect of Hct was investigated based on response current 5 seconds after beginning blood supply. The results are shown inFIG. 14.

Discussion of Experimental Results

As can be seen fromFIGS. 11 and 12, the response current tended to converge more quickly with the biosensor of Example 1 than with the biosensor of Comparative Example 1 when blood of different Hct was measured. First, as can be understood from looking at the 5-second values for response current for example, the difference in response current between Hct 20% and Hct 69% was smaller with the biosensor of Example 1 than with the biosensor of Comparative Example 1, and second, the response current for each sample had become uniform after about 8 seconds with the biosensor of Example 1. With the biosensor of Comparative Example 1, by contrast, it took about 15 seconds for the response current to become uniform.

These results mean that the biosensor of Example 1 is capable of performing proper concentration measurement in less time than the biosensor of Comparative Example 1. Moreover, because in the biosensors of Example 1 and Comparative Example 1 the internal channel60of capillary6has a small volume of 0.5 μL, it appears that the biosensor of Example 1 can precisely measure tiny quantities of blood.

Although reproducibility is less with the biosensor of Example 1 than with the biosensor of Comparative Example 1, this is probably attributable to variation in the formation of water-absorbing layer51because the coating material was applied by hand to form water-absorbing layer51. Consequently, reproducibility could probably be improved if water-absorbing layer51could be formed uniformly.

As can be seen fromFIGS. 13 and 14, the biosensor of Example 1 had a bias of +5% in response current 5 seconds after initiation of blood supply within the range of Hct 20–69%, while the biosensor of Comparative Example 1 had a bias of ±20%. This means that in the biosensor of Example 1 Hct affects response current less than it does in the biosensor of Comparative Example 1. These results show that the effect of blood Hct is reduced by the inclusion of a water-absorbing layer51such as that of the biosensor in Example 1.

As explained above, with the measuring instrument of the present invention it is possible to precisely measure concentration in a small quantity of sample liquid while controlling the effect of the solid components in the sample liquid and keeping the measurement time short.