Abstract:
A biosensor, a measuring instrument for biosensor, and a method of quantifying a substrate contained in specimen liquid to reduce a measurement error by the biosensor; a biosensor, comprising at least a pair of electrodes formed on an insulation substrate; a measuring instrument, comprising a supporting part for detachably supporting the biosensor, a plurality of connection terminals electrically connected to the electrodes, and a drive power supply for applying a voltage to the electrodes through the connection terminals; wherein the biosensor is inserted into the measuring instrument, either one of the electrodes of the biosensor is connected to first and second connection terminals fined to the measuring instrument only when the biosensor is inserted into the support part of the measuring instrument in a specified direction and has an electrode structure to allow the first connection terminal to conduct with the second connection terminal when a voltage is applied by the drive power supply.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application is a U.S. national phase of PCT Application No. PCT/JP01/10525, filed Nov. 30, 2001, and claims priority to Japanese Application No. 2000-364225, filed Nov. 30, 2000 and Japanese Application No. 2001-357144, filed Nov. 22, 2001. 
     TECHNICAL FIELD 
     The present invention relates to a biosensor for measuring the quantity of a substrate included in sample liquid and a measuring device for the biosensor. Further, the present invention provides a novel measuring method which reduces measurement errors caused by a biosensor. 
     BACKGROUND ART 
     Biosensors measure the quantity of a substrate included in sample liquid. The sensors utilize molecular recognition capability of bio material such as germ, enzyme, antibody, DNA, RNA and the like, and uses the bio material as a molecular recognizing element. In other words, when the bio material recognizes an objective substrate, it reacts such that the germ breathes, emits light, consumes oxygen, or causes enzyme reaction. The biosensors utilize those reactions and measure the quantity of the substrate included in the sample liquid. Among the biosensors, enzyme sensors have been promoted to practical use. For instance, an enzyme sensor for glucose, lactic acid, cholesterol, and amino acid is used in medical measurement and food industry. The enzyme sensor reduces an electron carrier with an electron produced by the reaction between the substrate and the enzyme included in the sample liquid, i.e., specimen. A measuring device measures the reduced amount of the electron carrier electrochemically, so that quantative analysis of the specimen is carried out. 
     Various kinds of biosensors, such as the one discussed above, have been proposed. A conventional biosensor, biosensor Z, will be described hereinafter.  FIG. 16(   a ) shows a perspective exploded view of biosensor Z.  FIG. 16(   b ) shows a structure of an electrode formed at a tip of biosensor Z. A method of measuring a quantity of a substrate in a sample liquid will be described with reference to  FIG. 16(   b ). 
     First, biosensor Z is inserted into a measuring device. The measuring device applies a given voltage across counter electrode  1103   a  and measuring electrode  1103   b . Then the sample liquid is supplied to inlet  1106   b  of a sample supplying path. The sample liquid is sucked into the supplying path due to capillary phenomenon, and passes on counter electrode  1103   a , which is nearer to inlet  1106   b , and arrives at measuring electrode  1103   b . Then reagent layer  1105  starts dissolving. At this time, the measuring device detects an electrical change occurring between counter electrode  1103   a  and measuring electrode  1103   b , and starts measuring the quantity. The quantity of the substrate included in the sample liquid is thus measured. 
     Specifically, oxidoreductase and an electron acceptor retained in the reagent layer dissolve into the sample liquid, and enzyme reaction progresses between the substrate in the liquid. Then the electron acceptor is reduced. After the reaction finishes, the reduced electron acceptor is oxidized electrochemically. A concentration of the substrate can be measured using an oxidation current measured when the acceptor is oxidized. 
     However, the conventional biosensor Z has some problems to be solved. In particular, when the measuring device detects the electrical change in reagent layer  1105 , various factors influence measurement accuracy and sensitivity of the measuring device. 
     First, an incorrect operation by a user influences them. For instance: (1) After the user supplies the sample liquid to the sample supplying path, the user adds another sample liquid before the measuring device completes the measurement; (2) The user tries to measure the quantity with a biosensor which have been already used; (3) The user supplies the sample liquid to a incorrect place; (4) The user inserts the biosensor into the measuring device in a wrong direction; and (5) When supplying the sample liquid, the user fails to pinpoint an inlet of the sample supplying path, has the sample liquid attach to a surrounding area, and thus has the sample liquid not run into the path. Thus some ways have been desired to avoid those incorrect operations which influence the measurement accuracy. In particular, preventing aged users from the incorrect operations is required. 
     Second, characteristics of an object to be measured influence them. For instance, when a glucose concentration of human blood is measured with a biosensor, a viscosity of the blood may influence measurement accuracy. Hematocrit, which is generally known as an index of blood viscosity, indicates a volume percentage of erythrocyte included in the blood. Blood in a person who does not suffer from anemia includes 50-60 volume % of water and 40-50 volume % of erythrocyte. If suffering from renal anemia due to chronic renal failure, a person has blood have the volume percentage of hematocrit decrease to less than 15%. Appropriate treatment requires to restrain the influence to hematocrit in the blood for accurate measurement of glucose concentration in the blood of, e.g., a diabetic. 
     Third, a temperature around the measuring device influences them. Measuring devices available in the market for biosensors have been downsized so that users can carry it with them. Soon after moving into indoors from the outside, a user may try to measure the quantity. In this case, the measurement may start before a temperature in the measuring device becomes stable. A sharp change in temperature influences the oxidation current corresponding to a substrate concentration, and thus may lower the measurement accuracy. A body temperature of the user, upon being transmitted to the measuring device via, e.g., the user&#39;s hand, might influence the measurement accuracy. 
     The present invention thus aims to provide a biosensor being handled easily and having excellent measurement accuracy, a method of measuring quantity using the biosensor, and a measuring device using the biosensor. 
     SUMMARY OF THE INVENTION  
     For solving the above problems, a first aspect of the present invention provides a biosensor for measuring the quantity of a substrate included in sample liquid. The biosensor is inserted to a measuring device which includes a supporting section for supporting detachably a biosensor which is formed of at least a pair of electrodes on an insulating board, plural connecting terminals electrically connected to the electrodes respectively, and a driving power supply for applying a voltage to the electrodes via the connecting terminals. One of the electrodes of the biosensor is connected to first and second connecting terminals of the measuring device only when the biosensor is inserted into the supporting section of the measuring device in a given direction. Then, the one of the electrodes becomes conductive due to a voltage application by the driving power supply. The electrodes have such a structure discussed above. 
     A conductive layer may be formed on at least a part of the insulating board, and the conductive layer is divided by slits, thereby forming a counter electrode and a measuring electrode, and upon request, a detecting electrode may be also formed. 
     A second aspect of the present invention aims to provide a measuring device to be used with a biosensor, and to measures a quantity of a substrate included in sample liquid. The measuring device includes a supporting section for supporting detachably the biosensor including at least a pair of electrodes on an insulating board, plural connecting terminals electrically connected to the electrodes, respectively, and a driving power supply for applying a voltage to the electrodes via the connecting terminals. The measuring device includes first and second connecting terminals can be connected to either one of electrodes of the biosensor only when the biosensor is inserted into the supporting section in a given direction. Thus conductivity can be detected between the first and the second connecting terminals by applying a voltage from the driving power supply to the first and second terminals, respectively. 
     It is also possible that the measuring device may determine that the biosensor is not inserted in the given direction if the conductivity is not detected. It is also possible that the measuring device may include an output section which outputs the determination to outside when the device determines that the biosensor is not inserted in the given direction. 
     A third aspect of the present invention provides a method of measuring a quantity of a substrate included in sample liquid with a biosensor. The biosensor includes an electrode section including: a counter electrode, a measuring electrode, and a detecting electrode on at least a part of an insulating board; a sample supplying path for supplying the sample liquid to the electrode section; and a reagent layer for reacting on the sample liquid supplied via the sample supplying path. The biosensor is inserted into a measuring device which includes a supporting section for supporting detachably the biosensor, connecting terminals, and a driving power supply for applying a voltage to the electrode section. When the biosensor is inserted into the supporting section of the measuring device, the driving power supply applies a voltage to a first electrode group and a second electrode group. The first group is formed of the counter electrode and the measuring electrode, and the second group is formed of the detecting electrode and one of the counter electrode and the measuring electrode. 
     In the biosensor, the detecting electrode among the counter electrode, the measuring electrode, and the detecting electrode is disposed most downstream along the sample supplying path, i.e., from a sample inlet along the sample flowing direction. It may be determined whether or not the sample liquid is supplied sufficiently for the measurement depending on whether or not respective electric currents from the first and second electrode groups exceed respective given thresholds. 
     After the electric current from the first electrode group exceeds the given threshold, if the current from the second group does not exceed the given threshold within a predetermined period, it may be determined that the sample liquid is insufficient. In this case, the measuring device may output the determination to the outside. 
     After the electric current from the first electrode group exceeds the given threshold, if the current from the second group does not exceed the given threshold within the predetermined period, an operator may hold a measuring step in order to add the sample liquid. 
     In the sample supplying path of the biosensor, the detecting electrode among the counter electrode, the measuring electrode, and the detecting electrode is disposed most downstream along the sample flowing direction from the sample inlet. An air hole for accelerating the flowing of the sample liquid is formed downstream against the detecting electrode. If the electric current from the second electrode group exceeds the predetermined threshold before the first group, and if the current from the first group does not exceed the threshold within a given period, it may be determined that the sample liquid is sucked from the air hole by mistake. 
     A measured quantity of the substrate corresponding to electric current detected by the electrode section may be compensated according to a lapse of time since the current from the first electrode group exceeds the threshold until the current from the second electrode group exceeds the threshold. 
     The measuring device may include a memory storing measured data which shows correspondence between a quantity of the substrate included in the sample liquid and a current detected by the biosensor. The measuring device refers to the measured data, thereby determining the quantity of the substrate corresponding to the detected current. 
     After the sample liquid is supplied to the sample supplying path, reaction between the sample liquid and the reagent layer is incubated during a certain time, and the quantity of the substrate is then measured. In this case, the incubating time may vary according to a lapse of time since the current from the first electrode group exceeds the threshold until the current from the second group exceeds the threshold. The driving power supply may apply a voltage to the first group and the second group alternately at constant intervals. 
     A fourth aspect of the present invention aims to provide a method of measuring a quantity of a substrate. This method uses a biosensor including a reagent layer which reacts specifically on the substrate in sample liquid. The method also uses a measuring device for measuring the quantity of the substrate included in the sample liquid from a sample produced by the reaction between the sample liquid and the reagent layer. The measuring device includes a temperature measuring section for measuring a temperature in the reaction progress between the sample liquid and the reagent layer and a temperature compensation memory for storing plural compensation tables of measured data. The compensation tables are prepared for each temperature range. The measuring device selects a compensation table according to a temperature measured by the temperature measuring section, and calculates a compensation value responsive to a measured quantity of the substrate for compensation. 
     The biosensor may include an electrode section including a counter electrode and a measuring electrode which are disposed on at least a part of an insulating board. The measuring device may apply a voltage to the electrode section and detect an electric current from the electrodes. 
     A fifth aspect of the present invention aims to provide a method of measuring, with a measuring device, a quantity of a substrate included in sample liquid supplied to a biosensor. The measuring device includes a temperature measuring section for measuring a temperature inside the measuring device. The temperature measuring section detects a temperature change between a temperature measured before the measurement of the substrate quantity and a temperature at the measurement. According to the temperature change, the measuring device determines whether the substrate quantity is to be measured or not. 
     If the temperature change exceeds a given threshold, the measurement may be cancelled. A temperature prior to the measurement may be measured intermittently. 
     A sixth aspect of the present invention aims to provide a method of measuring a quantity of a substrate included in sample liquid with a biosensor and a measuring device. The biosensor includes an electrode section including a counter electrode, a measuring electrode, and a reagent layer for reacting on sample liquid supplied to the electrode section which are disposed on at least part of an insulating board. The measuring device includes a supporting section for detachably supporting the biosensor, connecting terminals, and a driving power supply for applying a voltage to the electrode section. The driving power supply applies a voltage to the electrode section, and an electric current from the electrode section is detected, thereby measuring the quantity of the substrate included in the sample liquid. The measuring device applies a first voltage during a first period to the electrode section of the biosensor supported by the supporting section. After this voltage application during the first period, the voltage application is halted during a standby period. After the standby period, a second voltage is applied to the electrode section during a second period, and the current from the electrode section is measured, thereby measuring the quantity of the substrate. The first voltage is greater than the second voltage. 
     A seventh aspect of the present invention aims to provide a biosensor including two boards which are bonded to each other for forming a sample supplying path for taking sample liquid between the boards. The sample liquid is poured into an opening at respective ends of the boards as an inlet. The respective ends of the boards are located at different places from each other in a plan view of the biosensor. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  shows a biosensor system in accordance with a first exemplary embodiment of the present invention. 
         FIG. 2  is an exploded perspective view of a biosensor in accordance with a first exemplary embodiment of the present invention. 
         FIG. 3  shows combinations of recognizing sections of the biosensor depending on the presence of slits in accordance with the first embodiment. 
         FIG. 4  shows structures of the biosensor and a measuring device in accordance with the first embodiment. 
         FIG. 5  is a flowchart illustrating processes of measuring a quantity of a substrate included in sample liquid by the biosensor and the measuring device. 
         FIG. 6  is a flowchart illustrating processes of measuring a quantity of a substrate included in sample liquid by the biosensor and the measuring device. 
         FIG. 7  is a flowchart illustrating steps of measuring a quantity of a substrate included in sample liquid by the biosensor and the measuring device. 
         FIG. 8  illustrates a relation between a delay time and a compensation coefficient for compensating a measured quantity of a substrate. 
         FIG. 9  shows a profile at a measurement pre-process. 
         FIG. 10  shows a relation between a blood viscosity, a reaction time of reactive reagent layer and blood, and a measurement sensitivity. 
         FIG. 11  shows a glucose concentration (mg/dl) measured by a conventional method and a measurement pre-process of the present invention. 
         FIG. 12  shows data CA of a calibration curve. 
         FIG. 13  shows temperature compensation tables. 
         FIG. 14  shows relations between a temperature measured and measurement dispersion at each concentration of a substrate. 
         FIG. 15  shows a temperature change in a measuring device. 
         FIG. 16  is an exploded perspective view of a conventional biosensor. 
         FIG. 17  shows an exploded perspective view and a sectional view of a biosensor in accordance with a second exemplary embodiment of the present invention. 
         FIG. 18  is an enlarged plan view illustrating a sample supplying path of the biosensor shown in  FIG. 17 . 
         FIG. 19  shows an exploded view and a sectional view of another example of the biosensor. 
         FIG. 20  is an enlarged plan view illustrating a sample supplying path of the biosensor. 
         FIG. 21  illustrates a test method of sucking blood by the biosensor. 
         FIG. 22  illustrates another test method of sucking blood by the biosensor. 
     
    
    
     DETAILED DESCRIPTION OF EXEMPLARY EMBODIMENTS 
     Exemplary embodiments of the present invention will be described hereinafter with reference to the accompanying drawings. The embodiments discussed here are only examples, and the present invention is not necessarily limited to these embodiments. 
     Exemplary Embodiment 1 
     The first embodiment will be demonstrated hereinafter with reference to the accompanying drawings.  FIG. 1  shows a biosensor system in accordance with the first embodiment of the present invention. Biosensor system  1  includes biosensor  30  and measuring device  10  having biosensor  30  mounted detachably thereto. Sample liquid is dripped on sample-drip point  30   a  located at a tip of biosensor  30 . A quantity of a substrate included in the dripped sample liquid is measured by measuring device  10 . 
     Measuring device  10  includes, for instance, supporting section  2  to which biosensor  30  is detachably mounted and display  11  which shows a measured quantity of the substrate included in the sample liquid dripped on sample-drip point  30   a.    
     To measure a quantity of a substrate included in sample liquid with biosensor system  1 , first, a user inserts biosensor  30  into measuring device  10 . Then the user drips the sample liquid on sample-drip point  30   a  while measuring device  10  applies a certain voltage to electrodes of biosensor  30 . The sample liquid dripped, upon being sucked into biosensor  30 , make a reagent layer start dissolving. Measuring device  10  detects an electrical change generated between the electrodes of biosensor  30 , then starts measuring the quantity of the substrate. 
     Biosensor system  1  in accordance with the first embodiment is suitable to processing human blood as a sample liquid among others, and measuring a quantity of glucose, lactic acid, cholesterol included in the human blood as a substrate. Measuring the quantity of the substrate included in human body fluid is very important for diagnosis and medical treatment for a specific physiological abnormality. In particular, a diabetic is required to monitor his glucose concentration in the blood frequently. 
     The following demonstration refers to measuring a quantity of glucose included in human blood. However, biosensor system  1  in accordance with the first embodiment can measure a quantity of lactic acid, cholesterol and other substrates by selecting an appropriate enzyme as well. 
     Next, components forming biosensor  30  will be described with reference to  FIG. 2 , an exploded perspective view of biosensor  30 . Insulating board  31  (hereinafter called simply “board”) is made of, e.g., polyethylene terephthalate. On a surface of board  31 , a conductive layer, which is made of a noble metal such as gold and palladium, or an electrically conductive substance such as carbon, is formed by screen printing or sputtering evaporation. The conductive layer may be formed on the entire or at least a part of the surface. Reference numeral  32  denotes an insulating board having air hole  33  formed at its center. Spacer  34  having a notch is disposed between boards  31  and  32 , so that board  32  is integrated to board  31 . 
     On board  31 , the conductive layer is divided by a plurality of slits into counter electrode  37 , measuring electrode  38 , and detecting electrode  39 . In detail, the conductive layer is divided by the following slits: substantially arc-shaped slit  40  formed on counter electrode  37 ; slits  41   a  and  41   c  formed vertically to a side of board  31 ; slits,  41   b,    41   d,  and  41   f  and V-shaped slit  41   e.  The slits form counter electrode  37 , measuring electrode  38  and detecting electrode  39 . Each electrode may can be formed on at least a part of board  31 . Measuring device  10  may connected to the electrodes with lead wires. 
     Spacer  34  is placed for covering counter electrode  37 , measuring electrode  38 , and detecting electrode  39  on board  31 . The notch shaped in a rectangular provided at a center in a front section of spacer  34  forms sample supplying path  35 . The sample liquid is dripped to inlet  30   a  of sample supplying path  35 . The sample liquid dripped to inlet  30   a  is sucked by capillary phenomenon in an approximately horizontal direction (along arrow AR in  FIG. 2 ) toward air hole  33 . 
     Reference numeral  36  denotes a reagent layer formed by applying reagent, which contains enzymes, electron acceptors, amino acid, sugar alcohol and the like, to portions of counter electrode  37 , measuring electrode  38  and detecting electrode  39 , the portions which are exposed from the notch of spacer  34 . 
     The enzymes may employ the following materials: glucose oxidase, lactate oxidase, cholesterol oxidase, cholesterol estrase, uricase, ascorbate acid oxidase, bilirubin oxidase, glucose dehydrogenase, lactate dehydrogenase. 
     The electron acceptor preferably employs ferricyanide kalium, however, may employ p-benzoquinone and its derivatives, phenacine methor sulphate, methylene blue, and pherocane and its derivatives. 
     In the biosensor system in accordance with the first embodiment, glucose oxidase is used as oxidoreductase retained in reagent layer  36 , and ferricyanide kalium is used as the electron acceptor in order to measure the glucose concentration in human blood. 
     The oxidoreductase and the electron acceptor dissolve in the sample liquid (human blood in this embodiment) which is sucked into the sample supplying path, and then the glucose, a substrate in the sample liquid, reacts with the oxidoreductase and the electron acceptor, and the enzyme reaction progresses. Then the electron acceptor is reduced, thus producing ferrocyanide (ferricyanide kalium in this embodiment). After the reaction, the reduced electron acceptor, upon being oxidized electrochemically, generates a current from which the glucose concentration is measured. This series of reactions progress mainly in an area covering slits  40 ,  41   e  and detecting electrode  39 . The current produced by the electrochemical change is read out through measuring electrode  38  and detecting electrode  39 . 
     Reference numeral  42  denotes a recognizing section for recognizing, with measuring device  10 , a type of biosensor  30  and a difference in output characteristics among production lots. Slits  41   g  and  41   h  are combined to portions of counter electrode  37  and detecting electrode  39  corresponding to recognizing section  42 . The slits enables measuring device  10  to recognize the differences in output characteristics electrically. 
       FIG. 3  shows combinations of slits depending on the presence of slits  41   g,    41   h  in recognizing section  42  of biosensor  30 .  FIG. 3  illustrates seven types of combinations. For instance,  FIG. 3(   a ) shows recognizing section  42  of biosensor  30  for measuring cholesterol. In this case, slits  41   g  and  41   h  are not formed. 
       FIGS. 3(   b ),  3 ( c ), and  3 ( d ) illustrate recognizing section  42  of biosensor  30  for measuring lactic acid. In  FIG. 3(   b ), slit  41   h  is provided only in counter electrode  37 , thereby forming compensating section  43 . In  FIG. 3(   c ), slit  41   g  is provided only to detecting electrode  39 , thereby forming compensating section  44 . In  FIG. 3(   d ), slits  41   h  and  41   g  are provided to counter electrode  37  and detecting electrode  39 , respectively, thereby forming compensating section  43  and  44 , respectively. Further,  FIGS. 3(   e ),  3 ( f ), and  3 ( g ) illustrate recognizing section  42  of biosensor  30  for measuring glucose. In  FIG. 3(   e ), slit  41   g  is provided only to detecting electrode  39 , and slit  41   d  is formed up to slit  41   g.  And thus compensating section  44  is integrally formed with measuring electrode  38 . In  FIG. 3(   f ), slit  41   h  is added to the section in  FIG. 3(   e ), thereby forming compensating section  43 . In  FIG. 3(   g ), slit  41   f  is formed up to slit  41   h  shown in  FIG. 3(   f ). Thus, correcting sections  43  and  44  are integrally formed with measuring electrode  38 . 
     As such, a conductive area between the electrodes can be varied depending on patterns of the slits in recognizing section  42 . This enables measuring device  10  to recognize the differences in output characteristics (concentrations of glucose, cholesterol, lactic acid) of biosensor  30  and errors depending on production lots. Data and a control program, since being changed appropriately to the substrate according to the recognition, enables the device to be expected in exact measurement. This allows a user not to input compensating data using a compensating chip, and prevents the user from incorrectly handling the device. This embodiment discloses the biosensor having three electrodes. However, a number of electrodes may change, and a biosensor may have at least a pair of electrodes. The patterns of the slits other than those shown in  FIG. 3  may be formed. 
     Next, a structure of measuring device  10  will be explained in detail.  FIG. 4  shows structures of biosensor  30  (top view) and measuring device  10 . In biosensor  30 , counter electrode  37 , measuring electrode  38  and detecting electrode  39  are arranged along a flowing direction of a sample from sample-drip point  30   a  where detecting electrode  39  is placed most downstream. Counter electrode  37  may be exchanged between measuring electrode  38  in the arrangement order. Measuring electrode  38  and detecting electrode  39  are spaced at a given distance by slits  41   c  and  41   e . Thus, the device can determine, from an electric current changing according to an electrical change of the substrate, whether enough quantity of the sample liquid is sucked securely or not. 
     In measuring device  10 , reference numerals  12 ,  13 ,  14 ,  15 ,  16  and  17  denote connectors connected to areas A, B, C, D, E and F, respectively, which are produced by dividing recognizing section  42  of biosensor  30  into six areas. The six areas are grouped such that the groups correspond to slits  41   d,    41   f  and slits  41   g,    41   h.  Area A corresponds to measuring electrode  38 , area C corresponds to detecting electrode  39 , and area E corresponds to measuring electrode  38 . Area A is integrally formed with area B, and areas D and F correspond to compensating sections  43  and  44  shown in  FIG. 3 , respectively. Switches  18 ,  19 ,  20 ,  21  and  22  are provided between respective connectors  13 ,  14 ,  15 ,  16 ,  17  and a grounding (meaning a constant voltage, not necessarily “0”V. This definition is applicable to this description hereinafter.) Voltage to be applied to respective electrodes can be controlled at the grounding. Connectors  13 ,  14 ,  15 ,  16  and  17  are connected in parallel to the grounding. Switches  18  to  22 , upon being turned on and off under control, select a necessary connector out of connectors  13  to  17  which is used for the measurement. 
     Reference numeral  23  denotes a current/voltage converter connected to connector  12 , for converting a current flowing between measuring electrode  38  and other electrodes into a voltage. Reference numeral  24  denotes an A/D converter connected to current/voltage converter  23 , for converting a voltage supplied from circuit  23  into a pulse. Reference numeral  25  denotes a CPU for controlling to turn on and off the switches and calculating a content of the substrate included in the sample liquid based on the pulse supplied from A/D converter  24 . Reference numeral  11  denotes an LCD for displaying measured data calculated by CPU  25 . Reference numerals  26  and  28  denote temperature measuring sections for measuring temperatures inside measuring device  10 . Temperature measuring sections  26  and  28  are connected in parallel to each other between connector  12  and current/voltage converter  23 . 
     In measuring device  10  in accordance with the first embodiment, voltages (mV) converted from the currents flowing between the electrodes of biosensor  30  are used for detecting changes of the currents. In other words, the voltages indicate the currents flowing between the electrodes. 
     An operations of biosensor  30  and measuring device  10  will be demonstrated with reference to  FIG. 5  through  FIG. 7 , for measuring a content of a substrate in sample liquid by a method with biosensor  30  according to this embodiment. 
     First, it is determined whether or not biosensor  30  is properly inserted into supporting section  2  of measuring device  10  (Step S 1 ). Specifically, this is determined with a switch (not shown) in a connector shown in  FIG. 4 . If biosensor  30  is properly inserted (step S 1 : Yes), conductivity between areas A and B is tested (step S 2 ). As shown in  FIG. 3 , measuring electrode  38  has no slit formed therein for insulating one electrode itself such as slits  41   h  and  41   g.  In measuring electrode  38 , areas A and B are connected to connectors  12  and  13 , respectively. Areas A and B thus become conductive to each other without failure when biosensor  30  is inserted into measuring device  10  in a direction (a predetermined direction) such that a conductive layer of biosensor  30  is oriented normally. 
     Therefore, conductivity between areas A and B is tested by turning on switch  18 , so that the front and back sides of biosensor  30  can be determined. If the conductivity between areas A and B is not detected (step S 2 : No), it is determined that biosensor  30  is inserted front-side back (reversely). Then the measuring process terminates due to an error of detecting the front and back sides (step S 3 ). The error, when being detected, is preferably displayed on display  11 , or noticed as an alarm sound from a speaker. These preparations prevent the user easily from dripping blood to biosensor  30  by mistake while biosensor  30  is inserted front-side back. 
     When the conductivity between areas A and B is detected (step S 2 : Yes), it is determined whether or not voltages detected between area A and area C and between area A and area E are greater than 5 mV (step S 4 ). Switches  19  and  21  are simultaneously turned on, thereby allowing areas C and E to be considered to be electrically unified. Then a voltage is detected between area A and area C or E for determining whether biosensor  30  inserted in step  1  is an used one or not. This is determined since a reaction between reagent layer  36  and glucose in the blood has progressed to probably enlarge the detected voltage if biosensor  30  is the used one. 
     If it is determined that the voltage detected between area A and areas C is greater than 5 mV (step S 4 , Yes), it is recognized that biosensor  30  which is used is inserted, and the measuring process terminates due to an error of an used sensor (step S 5 ). If being detected, the error of used sensor is preferably displayed on display  11 , or noticed to a user as an alarm sound from a speaker. This prevents the user easily from dripping blood to biosensor  30  by mistake while used biosensor  30  is inserted. 
     Next, when the voltage detected between area A and areas C, E is not greater than 5 mV (step S 4 : No), the patterns of the slits is recognized by recognizing section  42  of biosensor  30  which is detected to be inserted at step S 1 . According to the recognizing result, CPU  25  changes data and a program into appropriate ones for output characteristics of the sensor (steps S 6  to S 10 ). In the first embodiment, three patterns of the slits are available, as shown in  FIGS. 3(   e ),  3 ( f ), and  3 ( g ), for a blood-sugar-level sensor which measures a glucose concentration. Specifically, first, conductivity between areas A and D is tested (step S 6 ). Switch  20  is turned on, and then the conductivity between areas A and D is tested, so that it may be determined whether or not biosensor  30  is proper to measure a blood sugar level and not proper to measure a quantity of lactic acid or cholesterol. 
     If the conductivity between areas A and D is not detected (step S 6 : No), it is determined that biosensor  30  is incompatible with the blood-sugar-level sensor. Then the measuring process terminates (step S 7 ), and display  11  shows an error message, or a speaker sounds an alarm for the user. These prevent the user from recognizing a measurement as a glucose concentration by mistake. 
     If the conductivity between areas A and D is detected (step S 6 : Yes), the conductivity between areas A and F is tested (step  8 ). Switch  22  is turned on. Then the conductivity between areas A and F is tested, so that the device can recognize differences in output characteristics due to production lots of biosensors  30  proper to blood-sugar-level sensors. CPU  25  automatically changes data and programs to which output characteristics corresponding to production lots have been reflected. Thus the user does not need a compensating chip. As a result, the biosensor and the measuring device can be handled more easily, and a higher accuracy of measurement can be expected. 
     If conductivity between areas A and F is detected (step S 8 : Yes), biosensor  30  is defined as a type shown in  FIG. 3(   g ), and result I is stored in a memory (not shown) (step S 9 ). If the conductivity between areas A and F is not detected (step S 8 : No), biosensor  30  is defined as a type shown in  FIG. 3(   e ) or  FIG. 3(   f ), and result II is stored in the memory (not shown) (step S 10 ). 
     After the type of biosensor  30  is recognized, it is determined again whether the voltage detected between area A and areas C, E is greater than 5 mV or not (step S 11 ). Switches  19 ,  21  are simultaneously turned on for detecting a current between area A and areas C, E. Then it is determined whether or not a user drips the sample liquid on biosensor  30  before measuring device  10  is ready for measurement. This process not only prevents positively the user from using used biosensor  30 , but also detects that the sample liquid has been dripped by the user before the measurement is available. 
     If the voltage detected between area A and areas C, E is greater than 5 mV (step S 11 : Yes), it is determined, as a drip error, that the sample liquid is dripped before the measurement is prepared. When being detected, the drip error is preferably displayed on display  11 , notified to a user with an alarm sound from a speaker, or displayed with LEDs (not shown) to give the user an alarm. The user can positively avoid a failure in operation by these operations, and a high accuracy of measurement can be expected. 
     If the voltage detected between area A and areas C, E is not greater than 5 mV (step S 11 : No), it is determined that the sample liquid is not dripped before the measurement is prepared. Then a completion of the preparation is notified to the user with LEDs (step S 13 ). When being detected, the error is preferably displayed on display  11 , notified to the user with an alarm sound from a speaker, or displayed with LEDs. Receiving this notice, the user takes blood as sample liquid from his body by himself and drips it to sample-drip point  30   a  of biosensor  30  inserted to measuring device  10 . 
     Next, it is determined whether or not enough quantity of the sample liquid is sucked through the sample supplying path from point  30   a  (steps S 14  to S 20 ). In biosensor  30 , counter electrode  37 , measuring electrode  38 , and detecting electrode  39  are arranged along sample supplying path  35  from sample-drip point  30   a  toward a downstream of the sample liquid flow. Detecting electrode  39  is placed most downstream. Either one of a group consisting of counter electrode  37  and measuring electrode  38 , or another group consisting of measuring electrode  38  and detecting electrode  39  is selected at a given interval. A voltage is applied to a selected group, so that it is determined whether or not the sample liquid is supplied in a quantity enough for measurement. In a conventional manner, a current change only between measuring electrode  38  and detecting electrode  39  is recognized. In the conventional manner, it is very difficult to identify a cause why the measurement does not start even though enough quantity of the sample liquid is supplied to the sample supplying path, or since the quantity is less than enough quantity for staffing the measurement. 
     Specifically, for the group of counter electrode  37  and measuring electrode  38 , switch  19  is turned off, and switch  21  is turned on for generating a voltage between areas A and E. For the group of measuring electrode  38  and detecting electrode  39 , switch  19  is turned on, and switch  21  is turned off for generating a voltage between areas A and C. As such, switches  19  and  21  are on-off controlled, thereby selecting and switching either one of the groups easily. For easy description, hereinafter, generating the voltage between counter electrode  37  and measuring electrode  38  is referred to as generating a voltage between areas A and E. Also generating a voltage between measuring electrode  38  and detecting electrode  39  is referred to as generating a voltage between areas A and C. 
     Further in this embodiment, as an example, a pair of areas A and E and a pair of areas A and C are switched every 0.2 seconds, and 0.2V is applied to each pair. It is determined whether or not respective voltages measured between areas A and E and between areas A and C reaches 10 mV (a given threshold). These numbers may be changed responsive to a type of biosensors. 
     Back to the flowchart shown in  FIG. 6 , the operations of the biosensor and the measuring device will be further described hereinafter. First, a voltage of 0.2V is produced between areas A and E which are located at the upstream portion of the sample supplying path, and it is determined whether or not the voltage measured between areas A and E exceeds 10 mV (step S 14 ). If the voltage measured does not exceed 10 mV (step S 14 : No), a voltage of 0.2V is applied between areas A and C located downstream of the path. Then it is determined whether the voltage measured between areas A and C exceeds 10 mV or not (step S 15 ). 
     If the voltage measured between areas A and C does not exceed 10 mV (step S 15 : No), it is determined whether or not 3 minutes have passed since the voltage was produced between areas A and E in step S 14  (step S 16 ). If the 3 minutes has not passed (step S 16 : No), the processes from step S 14  and onward are repeated. If respective voltages between areas A and E and between areas A and C do not reach 10 mV for 3 minutes (step S 16 : Yes), the measuring process terminates. 
     If the voltage between areas A and E is determined to reach 10 mV (step S 14 : Yes), it is determined whether or not the voltage between areas A and C reaches 10 mV (step S 17 ). If the voltage between areas A and C does not reach 10 mV (step S 17 : No), it is determined whether or not 10 seconds (a given period) have passed since the voltage between areas A and E was determined to reach 10 mV (step S 18 ). If the 10 seconds has not passed, the processes in steps S 17  and S 18  are repeated. While the 10 seconds passes, the measuring process temporarily halts until the voltage measured between areas A and C reaches 10 mV (step S 18 ; No). In this case, the sample liquid dripped is probably insufficient, it is preferable to display the error message on display  11 , or sound an alarm to a user from a speaker so that the user may understand that the sample liquid should be added. If the voltage measured between areas A and C does not reach 10 mV even after 10 seconds has passed (step S 18 : Yes), the measuring process terminates due to an error of specimen insufficient (step S 19 ). 
     While the 10 second passes since the voltage between areas A and E was determined to reach 10 mV in step S 14 , if a user adds the sample liquid, a final measurement accuracy is lowered. This was found by inventors. Specifically, while the user adds the sample liquid, the substrate in the sample liquid originally dripped has reacted on the enzyme included in reagent layer  36  and enzyme reaction has progressed. Thus a reduced substance has been produced before the measurement starts. After the added sample liquid reaches between areas A and C, the quantity of the substrate is possibly measured. In this case, the reduced substance already produced influences this measurement, i.e., makes the voltage apparently greater. In other words, as a time since the voltage between areas A and E is determined to reach 10 mV in step S 14  becomes longer, the measurement is influenced more by the reduced form. 
     In order to eliminate a measurement error caused by adding the sample liquid, a quantity of the substrate is compensated responsive to a measured voltage in measuring device  10  in accordance with this embodiment. The compensation depends on the lapse of time (delay time) since the voltage between areas A and E was determined to reach 10 mV in step S 14  until the voltage between areas A and C is determined to reach 10 mV in step S 17 . 
       FIG. 8  is a sensitivity compensation table illustrating a relation between the delay time and a compensation coefficient for the measured quantity of the substrate. The vertical axis represents the compensation coefficient, and the horizontal axis represents the delay time. For instance, if the delay time is 5 seconds, the measured quantity is compensated by 10% lower. As a result, 90% of the measured quantity becomes a compensated quantity. This kind of the sensitivity compensation table is stored in a memory (not shown). of measuring device  10 , and this table is referred when a final quantity of the substrate is calculated. 
     In biosensor  30  shown in  FIG. 2 , counter electrode  37  is formed such that slit  41   f  extends toward slit  41   c  and connects with slit  41   b.  Then a drip position error caused through dripping the sample liquid to air hole  33  by mistake, can be detected. In the flowchart shown in  FIG. 6 , if the voltage measured between areas A and C is determined to excess 10 mV not the voltage between areas A and E (step S 15 : Yes), it is determined, in 0.2 seconds after the determination, whether or not the voltage between areas A and E reaches 10 mV (step S 20 ). If the voltage between areas A and E does not excess 10 mV, it is determined that the sample liquid has been dripped to an incorrect position, and the measuring process terminates (step S 50 ). 
     If the sample liquid is normally dripped on sample-drip point  30   a , the liquid is sucked along sample supplying path  35  to air hole  33  and then moistens counter electrode  37 , measuring electrode  38  and detecting electrode  39  in this order. However, if the voltage measured only between areas A and C changes largely, a user has probably dripped the sample liquid to air hole  33  incorrectly. In this case, it is determined that an exact measurement is not expected, and the measuring process compulsorily terminates due to an error of dripping at a incorrect position. This can avoid a measurement error due to an incorrect operation by the user. 
     If the voltage measured between areas A and C is determined to reach 10 mV (step S 17 : Yes), or if the voltage measured between areas A and E is determined to reach 10 mV (step S 20 : Yes), enough quantity of the sample liquid is determined to be dripped. Then a pre-process for measuring the quantity of the substrate starts, and a timer (not shown) of measuring device  10  counts time (step S 21 ). 
     Next, conductivity between areas A and F is tested (step S 22 ). Switch  22  is turned on, and the conductivity is tested between areas A and F. If the conductivity is detected (step S 22 : Yes), it is determined whether result I identifying a type of biosensor  30  is stored in the memory in step S 9  or not (step S 23 ). If result I is stored (step S 23 : Yes), it is determined that the type of biosensor  30  is that shown in  FIG. 3(   g ). Calibration curve data is prepared using voltages measured when the reduced electron acceptor is oxidized electrochemically. Then calibration curve F 7  is prepared as the calibration curve data for specifying a concentration of the glucose in the sample liquid (step S 24 ). 
     On the other hand, when result II is stored (step S 23 : No), it is determined that the type of biosensor  30  is that shown in  FIG. 3(   e ), and calibration curve F 5  is prepared as the calibration curve data (step S 25 ). If the conductivity is not detected between areas A and F (step S 22 : No), it is determined that the type of biosensor  30  is that shown in  FIG. 3(   f ), and calibration curve F 6  is prepared as the calibration curve data (step S 26 ). 
     As discussed above, a difference in output characteristics of biosensor  30  is automatically recognized responsive to the slits in recognizing section  42  of biosensor  30 . The calibration curve data appropriate to the output characteristics is then automatically selected and set. Therefore, a user does not need a compensating chip, and CPU  25  switches automatically the calibration curve data to which the output characteristics depending on production lots are reflected. As a result, an incorrect measurement using user&#39;s incorrect data can be avoided, and a highly accurate measurement can be expected. 
     After the calibration curve is prepared in steps S 24  to S 26 , the measuring pre-process starts (step S 27 -S 29 ). The pre-process will be demonstrated with reference to  FIG. 9 , which illustrates a profile of the pre-process in accordance with the first embodiment. 
     In the profile shown in  FIG. 9 , the pre-process starts at time t 0 . Specifically, time t 0  is the time when the timer (not shown) of measuring device  10  starts counting time. The profile of the pre-process includes three consecutive periods, for instance, a first voltage period t 0 -t 1 , a standby period t 1 -t 2 , and a second voltage period t 2 -t 3 . 
     During the first voltage period, voltage V 1  is applied to areas A, C and E, to have the enzyme reaction progress. This increases a voltage measured by oxidizing ferrocyanide electrochemically similar to an exponential function. Next, during the standby period, voltage V 1  applied during the first voltage period is set at zero, and thus the ferrocyanide is not oxidized electrochemically, but the enzyme reaction keeps progressing. The ferrocyanide is thus accumulated. During the second voltage period, voltage V 2  is applied to areas A, C and E to oxidize the ferrocyanide accumulated during the standby period all at once. Then a quantity of discharged electron increases, and a high response current is thus observed at time t 2 . A current reaching the high response current decreases, as time passes, into a stable value i 3  at time t 3 . In the pre-process, switches  19  and  21  are simultaneously turned on in measuring device  10 , so that a voltage is applied to counter electrode  37  and detecting electrode  39  as one unit. 
     Recently, shortening a measurement time has been desired to upgrade a performance of the biosensor. When a quantity of a substrate is measured with a biosensor, a viscosity of the sample liquid critically influences measurement accuracy. This was found by the inventor. In particular, when human blood is measured as the sample liquid, blood with high viscosity (high Hct) lowers measurement sensitivity, and blood with low viscosity (low Hct) increases the measurement sensitivity. This phenomenon derives from a dissolving speed of a reagent layer in the blood, i.e., slow dissolution in high Hct and quick dissolution in low Hct. Thus the viscosity influences the measurement sensitivity of the biosensor. 
       FIG. 10  shows a relation between a blood viscosity, a reaction time of reactive reagent layer on the blood, and a measurement sensitivity. Data shown in  FIG. 10  is measured by a conventional method, which applies a voltage within a period corresponding to the second voltage period shown in  FIG. 9  and measures the voltage. As shown in  FIG. 10 , influence due to differences in viscosity (Hct in the case of blood) to measurement sensitivity increases at a shorter reaction time. Great difference is observed between the high Hct and the low Hct particularly at a reaction time around 5 sec. 
     Therefore, the conventional method tends to reveal a measurement error obviously due to blood viscosity. 
     During the first voltage period of the pre-process, reaction products produced at an initial stage of dissolving reagent layer  36  is thus compulsorily consumed by applying voltage V 1 . During the first voltage period, since the low Hct has a higher speed in enzyme reaction than the high Hct, greater reaction products are produced in the low Hct and thus greater reaction products are consumed. However, if a voltage is applied for too long period, reaction products are consumed too much, and responsivity of a voltage detected in the second voltage period may probably decline. Therefore, an effective first voltage period t 1 -t 0  may be 3 to 13 seconds. The voltage to be applied may be further increased, so that a voltage application time is preferably 2 to 10 seconds. Voltage V 1  may range preferably from 0.1 to 0.8V. 
     Next, during the standby period, the enzyme reaction progresses again, and the reaction products in the low Hct blood, the reaction products which have been consumed in the first voltage period, are quickly recovered and accumulated in approximately the same quantity as those in the high Hct blood. Too long a standby period or too short a standby period influences the final measurement sensitivity in a different way. 
     If the standby period is too short, a response value i 3  measured at time t 3  becomes too low, and a measurement error becomes great. If the standby period is too long, a difference in enzyme reaction speed between the low Hct blood and the high Hct blood probably becomes greater. The standby period is determined so that the difference in enzyme reaction speeds may not become greater. As a result, the standby period t 2 -t 1  is 1 to 10 seconds and preferably 2 to 10 seconds. 
     During the second voltage period, voltage V 2  starts being applied at time t 2 . And just after time t 2 , the voltage is not stable and requires a time to be stable. A voltage similar to that during the first voltage period is not necessarily applied, and a lower voltage than voltage V 1  is preferably applied. The lower voltage may be low enough to oxidize ferrocyanide kalium. The second voltage period t 3 -t 2  is thus preferably 2 to 10 seconds. Voltage V 2  is preferably 0.05 to 0.6V. Finally, value i 3  measured between areas A, C and area E at time t 3  is read out, and the quantity of the substrate (glucose) in the sample liquid is calculated. 
     The set time discussed above is particularly suitable for a quantity measuring with the biosensor including electrodes made of noble metal such as palladium. A reagent is not limited to glucose oxidase and/or glucose dehydrogenase and ferricyanide kalium, but includes amino acid, sugar alcohol. The set time is also suitable to a biosensor including organic acid. 
     After the sample liquid is supplied to sample supplying path  35 , the reaction of reagent layer  36  in the sample liquid is incubated in a certain period before the quantity of the substrate is measured. The incubate period may change depending on the laps of time since the voltage measured between areas A and E exceeds a threshold (10 mV) in step S 14  until the voltage between areas A and C exceeds the threshold (10 mV) in step S 17 . 
       FIG. 11  shows glucose concentrations (mg/dl) measured by the conventional method and the measuring pre-process discussed above for three types of blood having contents of hematocrit (Hct) of 25%, 45% and 65%. Reference mark R in  FIG. 11  denotes the measurement result by the pre-process. The other two results were measured by the conventional method with 15 seconds and 30 seconds of the reaction time. The pre-process was performed under the following condition: the first voltage period was 6 seconds; voltage V 1  was 0.5V; the standby period was 6 seconds; the second voltage period was 3 seconds; and voltage V 2  was 0.2V. As compared with a measurement for: the Hct was 45%; and the glucose concentration was 100 mg/dl, an actual measurement proved that low Hct (25%) blood and high Hct (65%) blood produce greater dispersion in the measurement, and the response values of low Hct disperse in a higher range and those of high Hct disperse in a lower range. 
     Further, the dispersion becomes greater at a shorter reaction time. At a reaction time of 15 seconds, the dispersion is produced by 10% higher (low Hct of 25%) and by 10% lower (high Hct of 65%). At a reaction time of 30 seconds, the dispersion is produced by 5% higher (low Hct of 25%) and by 5% lower (high Hct of 65%). In this pre-process, the dispersion is produced by 3% higher (low Hct of 25%) and by 3% lower (high Hct of 65%). At a reaction time of 15 seconds,  FIG. 11  teaches that the pre-process can reduce the dispersion due to the types of Hct while the reaction time is the same as that in the conventional method. 
     Back to  FIG. 7  again, the description of the measuring process continues hereinafter. The measuring pre-process starts, and 0.5V is applied between areas A and C, and between areas A and E for 6 seconds in the first voltage period (step S 27 ). After the first voltage period, the standby period is taken for 6 sec., and the voltage applied is cancelled in the standby period (step S 28 ). After the standby period, the second voltage period starts, and 0.2V is applied between areas A and C, and between areas A and E for 3 sec. (step S 29 ). Then value i 3  is read out (step  30 ). 
     After value i 3  is read out in step S 30 , temperature measuring sections  26  and  28  and switches  27 ,  29  disposed in measuring device  10  are controlled to measure a temperature in measuring device  10  (step S 31 ). Specifically, switch  27  is turned on, and measuring section  26  measures the temperature (step S 31 ). Then switch  27  is turned off, switch  29  is turned on, and measuring section  28  measures the temperature (step S 32 ). 
     The two temperatures measured by temperature measuring section  26  and  28  are compared with each other, and it is determined whether or not the difference between the two temperatures ranges within a given threshold (step S 33 ). If the difference is out of the threshold, the measuring process terminates due to a failure of either one of measuring section  26  or  28  (step S 33 : No). As such, plural temperature-measuring sections ( 26 ,  28 ) are disposed in measuring device  10 , and their measuring results are compared, so that a failure can be detected exactly and easily. This can avoid a measurement error caused by a measurement at an irregular temperature. The temperatures are measured just after the value has been read out in step S 30 ; however, the temperatures may be measured, for instance, when the pre-process starts in step S 21 . 
     If the difference between the two temperatures measured ranges within the given threshold (step S 33 : Yes), the temperatures are temporarily stored in a memory (not shown). At this time, the temperature measured by either one of sections  26  or  28  may be selected and stored, and the average of the two temperatures may be stored. Then a calibration curve, which should refer to value i3 measured in step S 30 , is specified (step S 34 ). The calibration curves prepared in steps S 24 , S 25  and S 26  are referred. If biosensor  30  corresponds to step S 24 , calibration curve F 7  is referred (step S 35 ). In the same manner, if biosensor  30  corresponds to step S 25 , calibration curve F 5  is referred (step S 36 ). If biosensor  30  corresponds to step S 26 , calibration curve F 6  is referred (step S 37 ). 
       FIG. 12  shows calibration curve data CA measured in steps S 34 , S 35  and S 36 . In data CA, a voltage measured in step S 30  and a concentration (mg/dl) of a substrate included in sample liquid are determined depending on each output characteristic F 1  to F 7  of biosensor  30 . For instance, if a measured voltage is 25 mV, and the biosensor corresponds to calibration curve F 5 , a substrate concentration of 14 (mg/dl) is stored in the memory. 
     Next, a concentration of the substrate selected in step S 35 , S 36  or S 37  is compensated by a compensation coefficient corresponding to the delay time which has been found in steps S 14  and S 17  and stored in the memory (step S 38 ). Specifically, the concentration is compensated by the following equation (1):
 
 D 1=(concentration of substrate)×[{100−(sensitivity compensation coefficient)}/100]
 
where D 1  is a compensated concentration of the substrate. This compensation eliminates a measurement error due to adding sample liquid by a user.
 
     Next, the concentration compensated in step S 38  is compensated according to the temperatures measured in steps S 31  to S 33  (step S 39 ). The temperature stored in the memory in step S 33  is read out, and a temperature compensation table shown in  FIG. 13  is referred, thereby determining a temperature compensation coefficient to be applied to concentration D 1 . 
       FIG. 13  shows temperature compensation tables. Compensation table T 10  is used for the temperature of 10° C. In the same manner, table T 15  is for the temperature of 15° C., and table T 20  is for the temperature of 20° C. The compensation tables specifies a relation between substrate concentration D 1  in the sample liquid and a temperature compensation coefficient is specified. The temperature compensation coefficient is determined based on a concentration at 25° C. as a reference, and shows a coefficient for compensation with respect to the concentration. Specifically, the compensation for temperature is performed according to the following equation (2):
   D 2 =D 1×(100−Co)/100 
where D 2  is a compensated concentration, D 1  is the concentration calculated in step S 38 , and Co is the temperature compensation coefficient specified by referring to the temperature compensation table.
 
     The inventors found experimentally that measurement accuracy was influenced by a combination of a measured temperature and a concentration of a substrate. The influence will be described hereinafter.  FIG. 14  shows relations between the measured temperature and measurement dispersion (bias) at each concentration of glucose. The measurement dispersion in  FIG. 14  is defined by a coefficient of a change of a concentration of glucose measured at 25° C. according to a change of the measured temperature.  FIG. 14(   a ) shows a relation between the dispersion and the measured temperature in the case of glucose concentration of 50 mg/dl at 25° C. Similarly,  FIG. 14(   b ) shows the relation for the glucose concentration of 100 mg/dl and the temperature of 25° C.  FIG. 14(   c ) shows the relation for the glucose concentration of 200 mg/dl and the temperature of 25° C.  FIG. 14(   d ) shows the relation for the glucose concentration of 300 mg/dl and the temperature of 25° C.  FIG. 14(   e ) shows the relation for the glucose concentration of 420 mg/dl and the temperature of 25° C.  FIG. 14(   f ) shows the relation for the glucose concentration 550 mg/dl and the temperature of 25° C. 
     These experimental data point out the following two tendencies. First, for the same glucose concentration, the measuring dispersion increases as a difference between a measured temperature and reference temperature 25° C. becomes greater. In detail, the dispersion increases in a negative direction as a measured temperature decreases from the reference temperature, and the dispersion increases in a positive direction as a measured temperature rises from the reference temperature. Second, the dispersion converges at the glucose concentration of 300 mg/dl, which seems a boundary, even though the glucose concentration increases. Specifically,  FIG. 14(   a ) indicates the dispersion of approximately 28% at 40° C.,  FIG. 14(   c ) indicates approximately 50%,  FIG. 14(   d ) indicates approximately 60%, and  FIG. 14(   f ) indicates approximately 50%. A similar tendency is found in a low temperature range such as a measured temperature of 10° C. 
     This tendency is reflected to the tables shown in  FIG. 13 . First, the measuring dispersion increases as a difference between a measured temperature and reference temperature of 25° C. becomes greater for the same glucose concentration. Second, the dispersion starts converging at the glucose concentration of 300 mg/dl as a boundary even though the glucose concentration increases. These two aspects are taken into consideration for preparing the tables. The measurement accuracy is remarkably improved by compensating a concentration referring to the temperature compensation table, in which combinations of measured temperatures and concentrations of the substrate are well considered, rather than compensating a concentration only based on a measured temperature. 
     In an operable temperature range of biosensor  30  (10° C. to 40° C. in this embodiment), a temperature compensation table for every 1° C. may be prepared, or the table for every given temperature range (e.g. 5° C.). If a temperature at a middle of the given temperature range is detected, a temperature compensation coefficient may be calculated by a linear interpolation with a temperature compensation table including the detected temperature. 
     Back to the flowchart in  FIG. 7 , concentration D 2 , which has undergone the temperature compensation discussed above, is output on display  11  of measuring device  10  as a final concentration of the substrate (step S 40 ). As discussed above, the time when the sample liquid is added, the measured temperature, and the combination of the measured temperature and the concentration are considered as influence factors to the measurement. A viscosity (Hct) of sample liquid is also considered as an influence factor. Those factors are taken into consideration when the quantity of a substrate is measured. As a result, the measurement accuracy is remarkably improved from the measurement by a conventional method. 
     The following method can be introduced in order to further decrease a measurement error due to temperature. 
     Before biosensor  30  is inserted into measuring device  10 , the temperature is measured successively and stored. After biosensor  30  is inserted, temperatures measured in steps S 31  and S 32  are compared with the stored ones. If large differences between the stored temperatures and the measured ones are found, the measuring process may compulsorily terminates due to a significant temperature change which influences a measurement error. 
     A portable biosensor system in accordance with this first embodiment, being carried easily, is exposed in various temperature changes depending on the outside environment. For instance, the biosensor system may be influenced by a temperature of a user&#39;s hand, or a sharp change in temperature when a user moves from outside to indoors. The sharp temperature-change can be expected, it takes reasonable time for measuring device  10  to be stabilized in its temperature change. 
       FIG. 15  shows temperature changes in measuring device  10 . A temperature change in device  10  moving from a place at a temperature of 10° C. to another place at a temperature of that of 25° C. is shown in  FIG. 15 . A temperature change in device  10  moving from a place at a temperature of 40° C. to a place of a temperature of 25° C. is also shown in  FIG. 15 .  FIG. 15  shows that it takes approximately 30 minutes to stabilize the temperature changes in an ambient temperature ranging from 10 to 40° C. If the temperature compensation is carried while the temperature changes, an exact temperature compensation may not be expected. 
     Therefore, if a great difference between the temperature stored in advance and the temperatures measured in steps S 31  and S 32 , the measuring process may compulsorily terminate due to the temperature change which may influence a measuring error. This further improves the accuracy of temperature compensation in measuring device  10 . A temperature may be measured before biosensor  30  is inserted into measuring device  10  at given intervals, e.g., 5-minute interval, or successively. Based on the magnitude of temperature change, the measuring process may be cancelled although a user tries to carry it out. 
     Exemplary Embodiment 2 
     The biosensor in accordance with the second exemplary embodiment will be demonstrated hereinafter. In this embodiment, an enzyme sensor is described. The sensor employs an enzyme as a molecule recognizing element which specifically reacts on a specific material contained in sample liquid. 
     An incorrect operation by a user influences a measuring accuracy. Thus the second embodiment discusses this problem. In particular, a user fails to drip sample liquid to an inlet of a sample supplying path, and the sample liquid attaches to a surrounding areas of the inlet. As a result, the sample supplying path cannot carry the sample liquid. Such kind of incorrect operations by a user may affect a measurement accuracy, and the ways how to avoid those mis-operations are demonstrated in this embodiment. 
     According to a structure shown in  FIG. 16  or  FIG. 2 , at the inlet, to which sample liquid is supplied, of the sample supplying path, an insulating board and a cover forming the path have respective ends of the same shape at the same location in a plan view. Therefore, a sample supplying angle becomes small. Or when the sample liquid attaches to a rear side (a side having no electrode formed thereon) of the insulating board by mistake, this sample liquid attached to the rear side may prevents the user from again supplying the sample liquid. As a result, the sample liquid is not supplied well, which causes a failure in measurement or a measurement error. 
     A biosensor which can accept the sample liquid exactly and easily will be specifically described hereinafter.  FIG. 17(   a ) is an exploded perspective view of the biosensor in accordance with the second embodiment.  FIG. 17(   b ) is a cross section at a center of the sample supplying path in the longitudinal direction of the biosensor. In  FIG. 17 , measuring electrode  52 , counter electrode  53  and detecting electrode  54  are formed on first insulating board  51 . Those electrodes are made of electrically conductive material. Detecting electrode  54  in this embodiment functions not only as an electrode for detecting a insufficiency of a specimen but also as a part of a reference electrode or as a part of the counter electrode. 
       FIG. 17  shows that the electrodes discussed above are disposed on the first insulating board; however, those electrodes may be divided and disposed also on second insulating board  58  to be a cover board located opposite to first board  51 . 
     Boards  51  and  58  are preferably made of polyethylene terephthalate, polycarbonate, polyimide or the like. 
     Each electrode is preferably made of electrically conductive material such as noble metal including gold, platinum, and palladium, or simple material such as carbon. They may be also made of composite material such as carbon paste or noble metal paste. In the former case, a conductive layer can be formed on board  51  or  58  easily by a sputtering evaporation method. In the latter case, a conductive layer can be formed on board  51  or  58  easily by a screen printing method. 
     The conductive layer is formed on an entire or a part of first insulating board  51  or second insulating board  58  by the sputtering evaporation method or the screen printing method. Then slits are provided by laser for forming and dividing the electrodes. The electrodes may be formed by the screen printing method or a sputtering evaporation method on a printed board or a masked board having electrode patterns formed in advance. 
     On the electrodes thus formed, reagent layer  55  is formed. Reagent layer  35  includes enzymes, electron carriers and hydrophilic high-polymer. The enzymes include glucose oxidase, lactate oxidase, cholesterol oxidase, cholesterol estrase, uricase, ascorbate acid oxidase, bilirubin oxidase, glucose dehydrogenase, lactate dehydrogenase. The electron carrier preferably employ ferricyanide kalium and may employ p-benzoquinone and its derivatives, phenacine methor sulphate, methylene blue, or pherocane and its derivatives. 
     The hydrophilic high-polymer employ, e.g. carboxymethyl cellulose, hdroxy-ethyl cellulose, hydroxy propyl cellulose, methyl cellulose, ethyl cellulose, ethyl hydroxyethyl cellulose, carboxy methyl ethyl cellulose, polyvinyl alcohol, polyvinyl pyrrolidone, polyamino acids such as poly-lysine, sulfonated polystyrene acid, gelatin and its derivatives, acrylic acid and its salts, methacrylic acid and its salts, starch and its derivatives, anhydrous maleci acid and its salts, or agarose gel and its derivatives. 
     First insulating board  51  and second insulating board  58  are bonded via spacer  56  in between for forming sample supplying path  57 , from which sample liquid is supplied. Spacer  56  has slit-shaped notch  57  formed therein. 
     A significant difference from the conventional biosensor is that first board  51  and second board  58  forming path  57  are placed with their ends at an inlet of sample supplying path  57  deviated each other and bonded. That is, respective ends are placed at different places from each other. This preparation is viewed from a plan view. In other words, first board  51  and second board  58  are in the same shape near the inlet of path  57 ; however, second board  58  and spacer  56  protrude toward the inlet with respect to first board  51 . 
     This allows the sample liquid to be sucked exactly and easily even though the sample supplying angle is small. This prevents the sample liquid from attaching to the rear side of first board  51 . Even if the sample liquid attaches to the rear side, the sample liquid can be supplied again smoothly. 
     The deviation of second board  58  from first board  51  at the ends thereof, that is, distance S 1  between points  63   a  and  64   a  is preferably not less than 0.1 mm and more preferably ranges from 0.25 to 1.0 mm, where center line L of path  57  shown in  FIG. 18  crosses with first board  51  and second board  58  at points  64   a  and  63   a,  respectively. 
     If being less than 0.1 mm, distance S 1  is too short. The sample liquid thus cannot be supplied well if the sample supplying angle is small as in the conventional biosensor. 
     If first board  51  has different shape from second board  58  near the inlet of path  57  as shown in  FIG. 19 , a similar advantage to that discussed above is measurable. In this case, the deviation at the ends thereof, i.e., center line L of path  57  shown in  FIG. 20  crosses with first board  51  at point  64   b  and crosses with second board  58  at point  63   b.  Distance S2 between point  63   b  and point  64   b  is preferably not less than 0.1 mm and more preferably ranging from 0.25 to 1.0 mm. 
     In the structures illustrated in  FIG. 17  to  FIG. 20 , a depth of the sample supplying path, i.e., a thickness of spacer  56 , ranges preferably from 0.05 to 0.3 mm in order to supply the sample liquid quickly to slit-shaped path  57 . 
     Spacer  56  is preferably made of polyethylene terephthalate, polycarbonate, polyimide, polybutylene terephthalate, polyamide, polyvinyl chloride, polyvinylidene chloride, or nylon. 
     For form sample supplying path  57 , first board  51  may be bonded to second board  58  integrated with spacer  56  into one unit. 
     Reagent layer  55  is disposed on entire or a part of a surface of the electrode, and however, may be disposed anywhere in sample supplying path  57  as long as it does not lower the performance of the biosensor. The sample liquid is supplied to the biosensor through path  57  having a structure discussed above by the capillary phenomenon. However, air hole  59  through which air flows outside the biosensor is necessary in path  57  in order to supply the sample liquid smoothly. Air hole  59  may shape in a rectangle, circle or polygon. 
     Air hole  59  may be located anywhere in path  57  as long as it does not block the supply of sample liquid. 
     Hydrophilic treatment which may be performed inside path  57  enables the sample liquid to be supplied into path  57  more quickly and accurately. The hydrophilic treatment is realized by developing surface active agent into or on second board  58 , or by roughing the surface of the board by sand-blasting, electric-discharge machining, non-glare process, mat process, or chemical plating. 
     In the biosensor discussed above, a current is generated by the reaction between a specific component in the sample liquid and reagent layer  55  containing enzymes. The current is conducted to an external measuring instrument (not shown) via lead-wires  60 ,  61 , and  62  of measuring electrode  52 , counter electrode  53 , detecting electrode  54  for being measured. 
     For the current measurement, a triple-electrode method employing measuring electrode  52 , counter electrode  53  and detecting electrode  54  is available as discussed in this embodiment. Besides the triple-electrode method, a double-electrode method employing only measuring electrode  52  and counter electrode  53  is available. Either method can produce the similar advantage to that of this embodiment; however, the triple-electrode method achieves more precise measurement. 
     EXAMPLE 1 
     A thin palladium film of 8 nm thickness was formed on the entire surface of the first insulating board made of polyethylene terephthalate by sputtering evaporation method. Then slits were provided on a part of the thin film by YAG laser, and thus the electrode was divided into a measuring electrode, a counter electrode and a detecting electrode. On top of that, water solution containing enzymes, electron carriers, and hydrophilic high-polymer was dripped such that the water solution covered the measuring electrode as a center and parts of the counter electrode as well as the detecting electrode. Then the water solution was dried to form a reagent layer. Further on top of that, a spacer made of polyethylene terephthalate and having a notch together with the second insulating board (cover) made of polyethylene terephthalate and having the air hole was bonded. As a result, the sample supplying path, i.e., a capillary which leads blood, was formed. 
     In order to confirm the advantage of the present invention, the following six types of blood-sugar value sensors having end-deviations (distance S) from the board to the spacer and cover were determined as: S=0 (a conventional sensor), 0.1, 0.25, 0.5, 1.0, and 2.0 mm. 
     Surface active agent is applied to the surface of the cover (inside of the sample supplying path) in order to supply the blood to the path more quickly.  FIG. 21  illustrates a test method for confirming a blood-sucking performance of the sensor depending on a blood-supplying angle in the blood-sugar value sensor discussed above. Table 1 shows the test result. 
                                                                                       TABLE 1                       S   Blood Supply                               (mm)   Angle (deg.)   1   2   3   4   5                                    Conventional   0   0   X   X   X   X   X       Sensor       15   Δ   X   X   Δ   X               30   Δ   Δ   Δ   X   Δ               45   ◯   ◯   ◯   ◯   Δ               90   ◯   ◯   ◯   ◯   ◯       Sensor According   0.1   0   ◯   Δ   Δ   ◯   Δ       to Present       15   ◯   Δ   ◯   ◯   Δ       Invention       30   ◯   ◯   ◯   ◯   Δ               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯           0.25   0   ◯   ◯   ◯   ◯   ◯               15   ◯   ◯   ◯   ◯   ◯               30   ◯   ◯   ◯   ◯   ◯               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯           0.5   0   ◯   ◯   ◯   ◯   ◯               15   ◯   ◯   ◯   ◯   ◯               30   ◯   ◯   ◯   ◯   ◯               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯           1.0   0   ◯   ◯   ◯   ◯   ◯               15   ◯   ◯   ◯   ◯   ◯               30   ◯   ◯   ◯   ◯   ◯               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯           2.0   0   ◯   ◯   ◯   ◯   ◯               15   ◯   ◯   ◯   ◯   ◯               30   ◯   ◯   ◯   ◯   ◯               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯               Definitions of the marks in the table:       ◯: The blood is sucked by one sucking.       Δ: The blood is sucked by two or three sucking operations.       X: The blood is not sucked at all.            
Definitions of the marks in the table:
         O: The blood is sucked by one sucking.   Δ: The blood is sucked by two or three sucking operations.   X: The blood is not sucked at all.       
     Table 1 tells that the conventional sensor having distance S=0 mm does not suck blood and requires several trials of supply for proper sucking when it has a small blood-supplying angle (0-30 degree). For a small blood-supplying angle, when a user supplies the blood to the sample supplying path, the blood attaches to the rear side of the insulating board firstly. Thus even if the user tries to supply the blood again, the blood is pulled by the blood attached to the rear side. This may be a reason why the conventional sensor does not work well. 
     The sensor of the present invention, on the other hand, sometimes requires several sucking operations when the blood-supplying angle is small even at the shortest distance S=0.1 mm; however the sensor does suck the blood at all. When the distance S is not less than 0.25 mm, the sensor sucks the blood easily at any sucking angle. 
       FIG. 22  illustrates a test method for testing the sensor in the blood-sucking performance depending on the blood-supplying angle. In this test, blood is attached to the rear side of the insulating board in advance at an area of 5 mm from the end of the board in order to prevent the blood from being sucked. Table 2 shows the test result. 
                                                                                       TABLE 2                       S   Blood Supply                               (mm)   Angle (deg.)   1   2   3   4   5                                    Conventional   0   0   X   X   X   X   X       Sensor       15   X   X   X   X   X               30   X   X   X   X   X               45   X   X   X   X   X               90   X   X   X   X   X       Sensor According   0.1   0   X   Δ   Δ   Δ   X       to Present       15   Δ   Δ   Δ   X   Δ       Invention       30   Δ   ◯   Δ   ◯   Δ               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯           0.25   0   ◯   ◯   ◯   ◯   ◯               15   ◯   ◯   ◯   ◯   ◯               30   ◯   ◯   ◯   ◯   ◯               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯           0.5   0   ◯   ◯   ◯   ◯   ◯               15   ◯   ◯   ◯   ◯   ◯               30   ◯   ◯   ◯   ◯   ◯               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯           1.0   0   ◯   ◯   ◯   ◯   ◯               15   ◯   ◯   ◯   ◯   ◯               30   ◯   ◯   ◯   ◯   ◯               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯           2.0   0   ◯   ◯   ◯   ◯   ◯               15   ◯   ◯   ◯   ◯   ◯               30   ◯   ◯   ◯   ◯   ◯               45   ◯   ◯   ◯   ◯   ◯               90   ◯   ◯   ◯   ◯   ◯               Definitions of the marks in the table:       ◯: The blood is sucked by one sucking.       Δ: The blood is sucked by two or three sucking operations.       X: The blood is not sucked at all.            
Definitions of the marks in the table:
         O: The blood is sucked by one sucking.   Δ: The blood is sucked by two or three sucking operations.   X: The blood is not sucked at all.       
     Table 2 shows that the conventional sensor having the distance S=0 mm can not suck the blood except the blood-supplying angle of 90 degree. On the other hand, the sensor of the present invention sometimes cannot suck the blood when distance S=0.1 mm at a small blood-supplying angle. However, the sensor can suck the blood easily at any blood-supplying angle when distance S is not less than 0.25 mm. 
     According to the second embodiment discussed above, the respective ends of first insulating board  51  and second insulating board  58  are deviated each other so that both the ends are placed at different places viewed from a plan view. This allows the sample liquid to be sucked exactly and easily. 
     In the second embodiment, an enzyme sensor as the biosensor is described. The present invention is similarly applicable to biosensors including a molecular recognition element reacting not only with the enzyme but also with germ, antibody, DNA, or RNA. 
     According to the sensor in accordance with the second embodiment, two boards bonded together form the sample supplying path, from which the sample liquid is taken out, between the boards. An opening is provided as an inlet at respective ends of both boards for accepting the sample liquid. The ends forming the inlet are located at different places or shaped in different forms viewed from a plan view of the biosensor. Thus the supply sample liquid can be sucked exactly and easily even if the sample-supplying angle is not enough (small). Further this prevents the sample liquid from attaching to the rear side of first insulating board  51 . If the sample liquid attaches to the rear side, a user can supply the sample liquid again to allowing the sample liquid to be supplied smoothly. 
     INDUSTRIAL APPLICABILITY 
     The present invention provides a biosensor which is handled by a user easily and exhibits an excellent measurement accuracy. The present invention also provides a measuring method using the biosensor as well as a measuring device using the biosensor. 
     REFERENCE NUMERALS 
     
         
           1  Biosensor System 
           2  Supporting Section 
           10  Measuring Device 
           11  Display 
           12  Connector 
           13  Connector 
           14  Connector 
           15  Connector 
           16  Connector 
           17  Connector 
           18  Switch 
           19  Switch 
           20  Switch 
           21  Switch 
           22  Switch 
           23  Current/Voltage Converter 
           24  A/D Converter 
           25  CPU 
           26  Temperature Measuring Section 
           27  Switch 
           28  Temperature Measuring Section 
           29  Switch 
           30  Biosensor 
           30   a  Sample-Drop Point 
           31  Insulating Board 
           32  Insulating Board 
           33  Air Hole 
           34  Spacer 
           35  Sample Supplying Path 
           36  Reagent Layer 
           37  Counter Electrode 
           38  Measuring Electrode 
           39  Detecting Electrode 
           40  Slit 
           41   a  Slit 
           41   b  Slit 
           41   c  Slit 
           41   d  Slit 
           41   e  Slit 
           41   f  Slit 
           41   g  Slit 
           41   h  Slit 
           42  Recognizing Section 
           43  Compensating Section 
           44  Compensating Section 
           51  First Insulating Board 
           52  Measuring Electrode 
           53  Counter Electrode 
           54  Detecting Electrode 
           55  Reagent Layer 
           56  Spacer 
           57  Sample Supplying Path 
           58  Second Insulating Board 
           59  Air Hole 
           60  Lead-Wire 
           61  Lead-Wire 
           62  Lead-Wire 
           63   a  Point 
           63   b  Point 
           64   a  Point 
           64   b  Point 
           1101  Insulating Board 
           1102   a  Counter Electrode Lead 
           1102   b  Measuring Electrode Lead 
           1103   a  Counter Electrode 
           1103   b  Measuring Electrode 
           1104  Resist 
           1105  Reagent Layer 
           1106  Spacer 
           1106   a  Sample Supplying Path 
           1106   b  Inlet of Sample Supplying Path 
           1107  Cover 
           1107   a  Air Hole