CA2521370C - Biosensor system - Google Patents
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- CA2521370C CA2521370C CA2521370A CA2521370A CA2521370C CA 2521370 C CA2521370 C CA 2521370C CA 2521370 A CA2521370 A CA 2521370A CA 2521370 A CA2521370 A CA 2521370A CA 2521370 C CA2521370 C CA 2521370C
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- analyte
- detection current
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
- G01N27/00—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means
- G01N27/26—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means by investigating electrochemical variables; by using electrolysis or electrophoresis
- G01N27/28—Electrolytic cell components
- G01N27/30—Electrodes, e.g. test electrodes; Half-cells
- G01N27/327—Biochemical electrodes, e.g. electrical or mechanical details for in vitro measurements
- G01N27/3271—Amperometric enzyme electrodes for analytes in body fluids, e.g. glucose in blood
- G01N27/3273—Devices therefor, e.g. test element readers, circuitry
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
- G01N33/00—Investigating or analysing materials by specific methods not covered by groups G01N1/00 - G01N31/00
- G01N33/48—Biological material, e.g. blood, urine; Haemocytometers
- G01N33/50—Chemical analysis of biological material, e.g. blood, urine; Testing involving biospecific ligand binding methods; Immunological testing
- G01N33/53—Immunoassay; Biospecific binding assay; Materials therefor
- G01N33/543—Immunoassay; Biospecific binding assay; Materials therefor with an insoluble carrier for immobilising immunochemicals
- G01N33/54366—Apparatus specially adapted for solid-phase testing
- G01N33/54373—Apparatus specially adapted for solid-phase testing involving physiochemical end-point determination, e.g. wave-guides, FETS, gratings
- G01N33/5438—Electrodes
Abstract
Biosensor system and method for the electrochemical determination of an analyte in a sample. An analyte reactant is mixed with the sample to form a reaction liquid which is contacted to electrodes for measurement of an analyte detection current. The analyte detection current is characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction including transfer of electrons through an electrode surface. A plurality of values of a curve of the analyte detection current versus time is measured and the analyte concentration is derived therefrom by means of an evaluation algorithm. Fast results of high accuracy are achieved in that the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously resulting in an analyte detection current versus time curve which comprises a section in which the current rises versus time and the analyte detection current is measured at at least two points of time in the rising section and the values of the analyte detection current resulting from said at least two measurements are used in the evaluation algorithm for temperature error compensation.
Description
Biosensor system The invention relates to a system and method for the electrochemical determination of an analyte in a blood sample. Such systems are often referred to as biosensor systems.
The concentration of a medically interesting analyte in blood is often determined by means of an electrochemical test. A highly specific reaction between the analyte and a test reagent, generally an enzyme, is performed which causes a change of the oxidation stage of the enzyme. The rate by which this change takes place depends on the concentration of the analyte. It is detected by means of a charge transfer to a solid electrode. The rate of charge transfer at the electrode is detected by applying a DC-voltage thereto and measuring the resulting DC-current. The concentration of the analyte is derived from this current measurement.
A typical reaction sequence of such an electrochemical test is shown in figure 1. An enzyme catalyzes an DOCSMTL: 2677481 \i oxidation of an analyte A thereby forming an analyte product P. Simultaneously the enzyme is reduced from its oxidized form Eo,, to its reduced form Ered. By this highly specific enzymatic reaction an electron is transferred from the analyte to the enzyme. The enzyme cycles back to the oxidized form thereby transferring an electron to a mediator which is transferred from its oxidized form M X
to its reduced form Mred= The mediator is reoxidized transferring an electron to a solid state electrode EL.
A very important example of the use of this test principle are glucose tests which are needed mainly to monitor the blood sugar status of diabetics. This is of utmost importance both medically and commercially, because the health status of diabetics highly depends on reliable, easily generated data on their blood glucose level. The enzymes used in glucose testing are generally glucose oxidase or glucose dehydrogenase. Prominent mediators are ferrocen, ferrocyanide and phenylendiamine.
The main field of application of the invention are such enzymatic amperometric tests, in particular for glucose.
It can, however, be applied to any amperometric test which includes a reaction sequence comprising the following redox reactions:
A) an analyte-specific redox reaction in a homogeneous (liquid) phase by which a product is formed which can, with application of a suitable voltage, be oxidized or reduced at a solid electrode;
B) a heterogeneous electrode reaction by which an electron is transferred from the product of the analyte-specific reaction to the electrode or from 35' the electrode to the product.
The concentration of a medically interesting analyte in blood is often determined by means of an electrochemical test. A highly specific reaction between the analyte and a test reagent, generally an enzyme, is performed which causes a change of the oxidation stage of the enzyme. The rate by which this change takes place depends on the concentration of the analyte. It is detected by means of a charge transfer to a solid electrode. The rate of charge transfer at the electrode is detected by applying a DC-voltage thereto and measuring the resulting DC-current. The concentration of the analyte is derived from this current measurement.
A typical reaction sequence of such an electrochemical test is shown in figure 1. An enzyme catalyzes an DOCSMTL: 2677481 \i oxidation of an analyte A thereby forming an analyte product P. Simultaneously the enzyme is reduced from its oxidized form Eo,, to its reduced form Ered. By this highly specific enzymatic reaction an electron is transferred from the analyte to the enzyme. The enzyme cycles back to the oxidized form thereby transferring an electron to a mediator which is transferred from its oxidized form M X
to its reduced form Mred= The mediator is reoxidized transferring an electron to a solid state electrode EL.
A very important example of the use of this test principle are glucose tests which are needed mainly to monitor the blood sugar status of diabetics. This is of utmost importance both medically and commercially, because the health status of diabetics highly depends on reliable, easily generated data on their blood glucose level. The enzymes used in glucose testing are generally glucose oxidase or glucose dehydrogenase. Prominent mediators are ferrocen, ferrocyanide and phenylendiamine.
The main field of application of the invention are such enzymatic amperometric tests, in particular for glucose.
It can, however, be applied to any amperometric test which includes a reaction sequence comprising the following redox reactions:
A) an analyte-specific redox reaction in a homogeneous (liquid) phase by which a product is formed which can, with application of a suitable voltage, be oxidized or reduced at a solid electrode;
B) a heterogeneous electrode reaction by which an electron is transferred from the product of the analyte-specific reaction to the electrode or from 35' the electrode to the product.
Evidently, each of these reactions may comprise a plurality of reaction stages. For example, the formation of Mred shown in figure 1 via the enzyme is a two-stage analyte-specific reaction. The mediator is, however, only needed if the direct charge transfer of the enzyme to the solid state electrode is too slow. Therefore, depending on the type of enzyme and electrode used, tests are also possible in which the mediator is omitted. On the other hand, the invention may be used in electrochemical tests comprising additional reaction stages.
In order to simplify the following description we refer hereafter only to the example of enzymatic tests in which reaction A ("forward reaction") is an enzymatic reaction which leads to the formation of a reduced mediator and reaction B ("reverse reaction") is a reoxidation of the mediator by transfer of electrons to the electrode. This should, however, not be understood as a limitation of the general applicability of the invention.
Electrochemical tests are produced in numerous different formats including electrodes suitable for immersion into the sample liquid. The invention is applicable to different formats, but is in particular directed to biosensor systems comprising disposable analysis elements and an evaluation instrument which is specifically adapted for evaluation thereof. Normally the components of the system are developed and supplied by the same manufacturer. The analysis elements contain the reagent system.required for a particular test (designated here "analyte reactant") and at least two electrodes. The evaluation instrument of the system generally has a holder for receiving a biosensor. When the biosensor is inserted into the holder an electrical contact is established between the electrodes of the biosensor and the instrument electronics.
The analyte reactant may be contained in a porous matrix such as a paper or porous plastic element which is in contact with the electrodes and to which the sample is applied. In recent years an alternative design of analysis elements has become increasingly important in which the analyte reactant and the electrodes are contained in a capillary space. The sample liquid is applied to an opening of the capillary space and drawn into the space by capillary action. Simultaneously the analyte reactant is dissolved and mixed with the sample liquid thereby forming a reaction liquid which is contacted to the electrodes for measuring of the analyte detection current which corresponds to the analyte concentration in the sample liquid. The invention refers particularly to biosensor systems comprising such a "capillary biosensor .
More details about such biosensor systems can be taken from the appropriate literature. In particular, a capillary biosensor system is described in WO 99/32881 and this document contains an extensive list of earlier publications from which a substantial amount of additional technical information can be taken.
Problems are caused by sources of error which have effects on the biosensor output. An important example are variations of the concentration of red blood cells, i.e.
hematocrit, in whole blood. This problem is discussed in the publication:
Tang et al.: "Effects of Different Hematocrit Levels on Glucose Measurement with Handheld Meters for Point-of-Care Testing", Arch Pathol Lab Med, 2000, pp. 1135 to 1140.
In order to simplify the following description we refer hereafter only to the example of enzymatic tests in which reaction A ("forward reaction") is an enzymatic reaction which leads to the formation of a reduced mediator and reaction B ("reverse reaction") is a reoxidation of the mediator by transfer of electrons to the electrode. This should, however, not be understood as a limitation of the general applicability of the invention.
Electrochemical tests are produced in numerous different formats including electrodes suitable for immersion into the sample liquid. The invention is applicable to different formats, but is in particular directed to biosensor systems comprising disposable analysis elements and an evaluation instrument which is specifically adapted for evaluation thereof. Normally the components of the system are developed and supplied by the same manufacturer. The analysis elements contain the reagent system.required for a particular test (designated here "analyte reactant") and at least two electrodes. The evaluation instrument of the system generally has a holder for receiving a biosensor. When the biosensor is inserted into the holder an electrical contact is established between the electrodes of the biosensor and the instrument electronics.
The analyte reactant may be contained in a porous matrix such as a paper or porous plastic element which is in contact with the electrodes and to which the sample is applied. In recent years an alternative design of analysis elements has become increasingly important in which the analyte reactant and the electrodes are contained in a capillary space. The sample liquid is applied to an opening of the capillary space and drawn into the space by capillary action. Simultaneously the analyte reactant is dissolved and mixed with the sample liquid thereby forming a reaction liquid which is contacted to the electrodes for measuring of the analyte detection current which corresponds to the analyte concentration in the sample liquid. The invention refers particularly to biosensor systems comprising such a "capillary biosensor .
More details about such biosensor systems can be taken from the appropriate literature. In particular, a capillary biosensor system is described in WO 99/32881 and this document contains an extensive list of earlier publications from which a substantial amount of additional technical information can be taken.
Problems are caused by sources of error which have effects on the biosensor output. An important example are variations of the concentration of red blood cells, i.e.
hematocrit, in whole blood. This problem is discussed in the publication:
Tang et al.: "Effects of Different Hematocrit Levels on Glucose Measurement with Handheld Meters for Point-of-Care Testing", Arch Pathol Lab Med, 2000, pp. 1135 to 1140.
The authors point out that even with the latest technology of biosensors large errors of the glucose value in the order of 20 to 30% are caused by variations of the hematocrit which may occur in practice. A number of possible mechanisms by which these errors may be caused are mentioned. It is noted that solutions to this problem are needed badly but no means for compensation of the hematocrit error are described.
Another important source of error are temperature variations. Amperometric tests generally show a strong dependence of the results from the temperature of the reaction liquid. Therefore, in some evaluation systems, the temperature of the reaction liquid is carefully controlled to a fixed value. In other systems the temperature is measured and a correction calculation is performed to compensate for temperature variations.
According to the above-mentioned WO 99/32881 an AC-measurement is made to achieve a correction for the combined effect of sample temperature and hematocrit. To this end an AC-voltage in the frequency range between about 2 kHz and about 10 kHz is applied to the electrodes, and the real and imaginary components of the impedance of the biosensor-sample system are determined.
The magnitude and the phase angle of the impedance are calculated therefrom and a look-up table stored in the instrument is consulted for a correction factor. This correction factor is applied to conventionally determined glucose values, thereby deriving a corrected glucose concentration.
Another important source of error are temperature variations. Amperometric tests generally show a strong dependence of the results from the temperature of the reaction liquid. Therefore, in some evaluation systems, the temperature of the reaction liquid is carefully controlled to a fixed value. In other systems the temperature is measured and a correction calculation is performed to compensate for temperature variations.
According to the above-mentioned WO 99/32881 an AC-measurement is made to achieve a correction for the combined effect of sample temperature and hematocrit. To this end an AC-voltage in the frequency range between about 2 kHz and about 10 kHz is applied to the electrodes, and the real and imaginary components of the impedance of the biosensor-sample system are determined.
The magnitude and the phase angle of the impedance are calculated therefrom and a look-up table stored in the instrument is consulted for a correction factor. This correction factor is applied to conventionally determined glucose values, thereby deriving a corrected glucose concentration.
In order to achieve a very good accuracy of the detected concentration while simultaneously getting very fast results the invention proposes a biosensor system for the electrochemical determination of an analyte in a sample, the system comprising an analyte reactant and at least two electrodes, wherein the analyte reactant is mixed with the sample liquid to form a reaction liquid and the reaction liquid is contacted to the electrodes for measurement, in response to a DC-voltage applied to the electrodes, of an analyte detection current which corresponds to the analyte concentration in the sample liquid, the analyte detection current being characteristic of the analyte concentration in the sample by virtue of a rea:ction sequence comprising an analyte 1.5 specific reaction taking place in the reaction liquid and an electrode reaction including transfer of electrons through an electrode surface and electronic circuitry comprising a DC-voltage source for applying the required DC-voltage to the electrodes and a measurement and evaluation electronics for measuring a plurality of values of a curve of the analyte detection current versus time over a plurality of measurement times and for deriving the analyte concentration therefrom, which is characterized in that the DC voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously resulting in an analyte detection current versus time curve which comprises a section in which the current rises versus time (as a result of the kinetics of the analyte specific reaction and the kinetics of the electrode reaction), the analyte detection current is measured at at least two points of time in the rising section and the at least two measurements are used in an algorithm for temperature error compensation. Also a corresponding method is proposed.
The inventors have found that an error compensation which is both fast and accurate can be achieved by using current measurement values from a very early stage of the analysis reaction.
Thus in one aspect of the invention there is provided a biosensor system for the electrochemical determination of an analyte concentration in a blood sample, the system comprising:
an analyte reactant and at least two electrodes, wherein the system is adapted for:
mixing of the analyte reactant with the blood sample to form a reaction liquid and contacting of the reaction liquid to the electrodes for measurement of an analyte detection current flowing between the electrodes in response to a DC-voltage applied thereto, the analyte detection current being characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction taking place at an electrode of said at least two electrodes including transfer of electrons through an electrode surface; and electronic circuitry comprising a DC-voltage source for applying the required DC-voltage to the electrodes and a measurement and evaluation electronics for measuring a plurality of values of a curve of the analyte detection current versus time over a plurality of ineasurement times and for deriving the analyte concentration therefrom by means of an evaluation algorithm characterized in that the electronic circuitry is adapted for:
7a application of the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously in accordance with a relationship defined by an analyte detection current versus time curve which comprises a rising section in which the analyte detection current rises versus time, measuring of the analyte detection current at at least two points of time in the rising section of the curve;
and using the values of the analyte detection current resulting from said at least two measurements in the evaluation algorithm for a temperature error compensation.
In another aspect of the invention there is provided a method for the electrochemical determination of an analyte concentration in a blood sample, comprising the steps of:
mixing an analyte reactant with the sample to form a reaction liquid, contacting the reaction liquid to a pair of electrodes, measuring a plurality of values of a curve of an analyte detection current versus time over a plurality of measurement times, the analyte detection current flowing between the electrodes in response to a DC-voltage applied thereto and being characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction taking place at an electrode of said pair including transfer of electrons through an electrode surface; and DC)C'SMTL: 2677481 \ 1 7b deriving the analyte concentration by means of an evaluation algorithm comprising use of said plurality of values of the analyte detection current, characterized in that:
the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously resulting in the analyte detection current versus time curve which comprises a rising section in which the current rises versus time; and the analyte detection current is measured at at least two points of time in the rising section of the curve and the values of the analyte detection current resulting from said at least two measurements are used in the evaluation is algorithm for temperature error compensation.
The invention becomes better understandable from the following description in combination with the figures, wherein:
Fig. 1 shows a graphical representation of a typical analysis reaction sequence, Fig. 2 shows a graphical representation of DC-current versus time measured in a test according to the prior art, Fig. 3 shows a graphical representation of DC-current versus time measured in a test according to the invention, Fig. 4 shows a graphical representation corresponding to figure 3 for the same analyte concentration but a plurality of different temperatures and hematocrit values of the reaction liquid, I)OCSM"I'L: 2677481 \ I
7c Fig. 5 shows a graphical representation corresponding to figure 4 but with a normalized time axis, Fig. 6 shows a graphical representation of AC-current versus time for the same analyte concentration but a plurality of different temperatures and hematocrit values of the reaction liquid, Fig. 7 shows a graphical representation corresponding to figure 5 but with a linear transformation of the current values, DOC'SM"I'L: 2677481\l Fig. 8 shows the components of a typical biosensor system to which the invention refers in top view, Fig. 9 shows a schematic diagram of parts of a preferred electronic circuitry, Fig. 10 shows a combined view of DC current and AC
current measurements performed with the circuitry of figure 9.
Fig. 1 has been described before.
Fig. 2 is taken from US-patent 5,243,516 and shows a typical example of the timing sequence in prior art tests. After an analysis element has been plugged into the evaluation instrument a sample detection voltage is applied to the electrodes. No current flows as long as there is no sample in a sample well of the analysis element bridging the electrodes. However, as soon as the sample well is filled a current spike CS is sensed indicating that a drop of sample has been dosed into the analysis element and bridges the electrodes. This point of time is designated "dose detect" (DD). After current spike CS has been sensed the-sample detect voltage is removed from the electrodes and the analyte-specific forward reaction takes place during an incubation period IP. After the forward reaction has taken place an analysis voltage suitable for the electrode reaction (reverse reaction) is applied to the electrodes. Figure 2 shows typical shapes of the resulting functional relationship I(t) of the current I versus time for four different values of the glucose concentration as indicated in the figure. These curves of the DC-current versus time are hereafter designated "I(t)-traces".
The inventors have found that an error compensation which is both fast and accurate can be achieved by using current measurement values from a very early stage of the analysis reaction.
Thus in one aspect of the invention there is provided a biosensor system for the electrochemical determination of an analyte concentration in a blood sample, the system comprising:
an analyte reactant and at least two electrodes, wherein the system is adapted for:
mixing of the analyte reactant with the blood sample to form a reaction liquid and contacting of the reaction liquid to the electrodes for measurement of an analyte detection current flowing between the electrodes in response to a DC-voltage applied thereto, the analyte detection current being characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction taking place at an electrode of said at least two electrodes including transfer of electrons through an electrode surface; and electronic circuitry comprising a DC-voltage source for applying the required DC-voltage to the electrodes and a measurement and evaluation electronics for measuring a plurality of values of a curve of the analyte detection current versus time over a plurality of ineasurement times and for deriving the analyte concentration therefrom by means of an evaluation algorithm characterized in that the electronic circuitry is adapted for:
7a application of the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously in accordance with a relationship defined by an analyte detection current versus time curve which comprises a rising section in which the analyte detection current rises versus time, measuring of the analyte detection current at at least two points of time in the rising section of the curve;
and using the values of the analyte detection current resulting from said at least two measurements in the evaluation algorithm for a temperature error compensation.
In another aspect of the invention there is provided a method for the electrochemical determination of an analyte concentration in a blood sample, comprising the steps of:
mixing an analyte reactant with the sample to form a reaction liquid, contacting the reaction liquid to a pair of electrodes, measuring a plurality of values of a curve of an analyte detection current versus time over a plurality of measurement times, the analyte detection current flowing between the electrodes in response to a DC-voltage applied thereto and being characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction taking place at an electrode of said pair including transfer of electrons through an electrode surface; and DC)C'SMTL: 2677481 \ 1 7b deriving the analyte concentration by means of an evaluation algorithm comprising use of said plurality of values of the analyte detection current, characterized in that:
the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously resulting in the analyte detection current versus time curve which comprises a rising section in which the current rises versus time; and the analyte detection current is measured at at least two points of time in the rising section of the curve and the values of the analyte detection current resulting from said at least two measurements are used in the evaluation is algorithm for temperature error compensation.
The invention becomes better understandable from the following description in combination with the figures, wherein:
Fig. 1 shows a graphical representation of a typical analysis reaction sequence, Fig. 2 shows a graphical representation of DC-current versus time measured in a test according to the prior art, Fig. 3 shows a graphical representation of DC-current versus time measured in a test according to the invention, Fig. 4 shows a graphical representation corresponding to figure 3 for the same analyte concentration but a plurality of different temperatures and hematocrit values of the reaction liquid, I)OCSM"I'L: 2677481 \ I
7c Fig. 5 shows a graphical representation corresponding to figure 4 but with a normalized time axis, Fig. 6 shows a graphical representation of AC-current versus time for the same analyte concentration but a plurality of different temperatures and hematocrit values of the reaction liquid, Fig. 7 shows a graphical representation corresponding to figure 5 but with a linear transformation of the current values, DOC'SM"I'L: 2677481\l Fig. 8 shows the components of a typical biosensor system to which the invention refers in top view, Fig. 9 shows a schematic diagram of parts of a preferred electronic circuitry, Fig. 10 shows a combined view of DC current and AC
current measurements performed with the circuitry of figure 9.
Fig. 1 has been described before.
Fig. 2 is taken from US-patent 5,243,516 and shows a typical example of the timing sequence in prior art tests. After an analysis element has been plugged into the evaluation instrument a sample detection voltage is applied to the electrodes. No current flows as long as there is no sample in a sample well of the analysis element bridging the electrodes. However, as soon as the sample well is filled a current spike CS is sensed indicating that a drop of sample has been dosed into the analysis element and bridges the electrodes. This point of time is designated "dose detect" (DD). After current spike CS has been sensed the-sample detect voltage is removed from the electrodes and the analyte-specific forward reaction takes place during an incubation period IP. After the forward reaction has taken place an analysis voltage suitable for the electrode reaction (reverse reaction) is applied to the electrodes. Figure 2 shows typical shapes of the resulting functional relationship I(t) of the current I versus time for four different values of the glucose concentration as indicated in the figure. These curves of the DC-current versus time are hereafter designated "I(t)-traces".
If all required test conditions are met the shape of the I(t)-traces corresponds (after a surge time ST in which the I(t)-trace is substantially influenced by the particulars of the measurement electronic system) to a characteristic function proportional to 1/4t. A deviation from this "Cottrell current curve" indicates deviations from the required test conditions. US-patent 5,243,516 proposes to make a plurality of current measurements at a plurality of measurement times during the period in which the reverse electrode reaction takes place and to use a simple mathematical method to control whether the I(t)-traces as shown in figure 2 are in agreement with the Cottrell current. If this is not the case a malfunction of the system is assumed and indicated.
Figure 3 shows a typical shape of an I(t)-trace which results when the analysis voltage (DC-voltage required for the "reverse" electrode reaction) is applied not at the end of an incubation period required for the analyte specific forward reaction but rather at a much earlier point of time (preferably at "Dose Detect" or within at most 500 milliseconds, preferably at most 300 milliseconds and most preferably at most 100 milliseconds thereafter). Conveniently no special sample detect voltage is used but rather the, analysis voltage is applied before the sample is contacted to the electrodes.
The shape of the I(t)-trace shown in figure 3 is the result of a rather complicated combination of a plurality of influencing factors. In particular it depends on the speed of the enzymatic reaction which again depends not only on the concentration of the analyte but also on a plurality of potentially disturbing factors such as the temperature, and the diffusion of the reactants (which again is influenced by the temperature and the presence of particles, in particular blood cells, in the sample).
The product generated by the enzymatic reaction is simultaneously consumed by the electrode (reverse) reaction which is also influenced by a plurality of 5 factors, mainly by the concentration of the product of the enzymatic reaction in the vicinity of the electrode surface which again depends on the diffusion conditions in the sample liquid. In any case it is characteristic for the invention that the I(t)-traces include an initial 10 section in which the DC-current rises versus time. This section is designated rising section RS.
It has to be noted that an I(t)-trace measured in practice may have a more complicated shape, depending on the experimental setup. In particular it may not be , .monotonically rising in its section afterr application of the analysis voltage. Rather there may be an intermediate maximum as shown in dashed line in figure 3. Such a shape is, however, not due to the reaction sequence explained but rather due to transient behavior of the experimental setup, in particular of the measurement electronics. The rising section of the I(t)-trace as defined herein is the section preceding the maximum M which is due to the redox reaction sequence.
In the context of the invention it has been found that in spite of this complicated situation measurements of the analyte detection current which are made at a point of time within the rising section RS of the I(t)-trace (before it reaches its maximum M) can be used to improve the quality of the analytical result derived from the measurement. In particular by measuring the DC-current at at least two points of time in the rising section RS it is possible to compensate for errors caused by variations of the temperature of the reaction mixture. It is also il possible to derive the analyte concentration itself from current measurements in the rising section RS. Additional error corrections, in particular to compensate for errors caused by variations of the hematocrit values of the sample, are possible if also AC-measurements are performed during the rising section of the I(t)-trace.
Most preferably measurements of the DC-current and of the AC-current are performed simultaneously, i.e. there is an at least partial overlap of the measurement periods during which the DC- and AC-measurements are performed.
The invention achieves superior results regarding analytical tests at high speed and with very good analytical quality. In this context it has to be noted that generally enzymes providing a high specifity for a particular analyte react relatively slowly. With the classical method this increases the incubation period and therefore the time required for generating the analytical result. If enzymes are used which are optimized for short reaction times this generally requires a reduction of the specifity and thus*of the quality of the test. The invention allows to fulfill both requirements (high specifity and fast response) at the same time.
Figures 4 to 7 show results from experimental work performed in the context of the invention.
Figure 4 shows I(t)-traces as in figure 3, i.e. with the analysis voltage applied at dose detect, for a single glucose concentration (500mg/dl) but for varying values of the hematocrit and of the reaction liquid temperature.
The graph shows curves for three hematocrit values (20%, 45% and 70 %) and three temperatures (100, 25 and 40 C) each. The individual I(t)-traces are distinguished by using different symbols as shown in the column on the right side of the graphs. The same symbols are also used in figures 5 and 6. The data were derived from the original current measurements applying a baseline correction:
IBC (t) =I rg (t) -Cx (t-1/2x eo.1/t) wherein IBC . baseline corrected current Iorg . original Current C . empirical constant This baseline correction subtracts contributions caused by electrical charge carriers contained in the sample but not relating to the glucose concentration (t-1/2-term) and by the damping effect of the electronics (eo''/t-term) .
A further correction was necessary in view of the fact that it is not possible to prepare samples having exactly the desired values of the glucose concentration and of the hematocrit. Rather the samples had to be prepared by diluting available whole blood samples and adding glucose as required. Resulting minor deviations from the exact values shown in figure 4 were corrected by linear regression.
Figure 4 shows the following characteristics:
- Each of the curves includes a rising section essentially starting at dose detect (time 0) up to a maximum.
- The position of the maximum on the time scale depends on the temperature. The higher the temperature the earlier the maximum occurs (the drop of the I(t)-trace after the maximum is not visible for the 10 C-traces because it occurs after the maximum measurement time (5sec) shown). It is, however, possible to electronically detect a maximum with each of the traces).
- For each temperature the hematocrit causes a shift of the I(t)-trace: An increase of the hematocrit causes a decrease of the measured current values.
Figure 5 shows the same data as figure 4 but with a normalized time scale according to:
tnorm= t /tmax wherein tmax: time after dose detect at which the maximum of the respective 1(t) -trace occurs Figure 6 shows the results of AC-measurements with a frequency of 2 kHz versus time for the same samples as figure 4. The quantity shown is the admittance A
(reciprocal of impedance Z) measured in S(Ohm-1).. The curves show relatively strong initial variations which relate to the transient behavior of the measurement electronics. After at most two seconds, however, a stable value is achieved which shows pnly a very slow increase versus time which is due to a strong damping of the electronics used for the experiments.
Figure 7 shows the same data as figure 5 but after=a transformation of the current values according to:
Itf=lnorm x(1 + deltaZ x A) wherein deltaZ is calculated according to:
deltaZ= (Z-Zmed) wherein Z: impedance measured between the electrodes with an AC-voltage of 2 kHz at a point of time after the initial Z variations (e.g. at 3 sec in figure 6).
Zmed: Median of the Z values at this point of time A: Empirically determined weighing factor for optimum coincidence of the curves shown in figure 7.
The series of figures 4 to 7 shows that by relatively simple mathematical steps the complicated shape of the I(t)-traces shown in figure 4 can be reduced to essentially a single coinciding curve as shown in figure 7. In other words the I(t)-traces shown in figure 4 having a non-linear dependence from temperature and hematocrit can be transformed to I(t)-traces having a linear dependence on these disturbing factors by using a suitable transformation of the time axis.
These results show that - deviating from the prior art -reliable results, including correction for such disturbing factors as temperature and hematocrit, can be derived from amperometric measurements in which the enzymatic forward reaction takes place simultaneously with the reverse electrode reaction. It is not only possible but even favorable to use current measurement values from the rising section of the I(t)-traces. This again allows to substantially reduce the required time for performing a measurement because it is not necessary to introduce an incubation period for first performing the forward reaction and it is not even necessary to wait until the forward reaction and the reverse reaction lead to a steady state (in which the mediator is generated with the same speed in which it is consumed, i.e. its concentration is constant).
A simple way of performing an analysis of a sample having 5 an unknown glucose concentration comprises the following steps:
- An I(t)-trace as shown in figure 3 is measured.
- A baseline correction is performed.
- tma,, is determined.
10 - A current value measured at a time which corresponds to a predetermined point of time on the normalized time axis (e.g. 0.7 x tma,) is used for further evaluation.
- This current value is corrected on the basis of an AC
ls measurement.
- The resulting corrected current is transformed to a glucose concentration based on a calibration performed for the same point of time on the normalized time axis (e.g. 0.7 x tma,) This, however, is only one of several possibilities to evaluate measurements made in accordance with the invention. Figures 4 to 7 show,that the shape of the rising section of the I(t)-traces is characteristic for the combination of the influencing factors including a "variance" (the concentration of the analyte, e.g.
glucose) and a plurality of "covariances" (disturbing factors such as temperature and hematocrit). Therefore the effects of the covariances can be separated and consequently eliminated by measuring a plurality of values of the DC-current in the rising section of the I(t)-trace and by feeding these values into a suitable algorithm.
In place of the described mathematical-analytical method for example numerical methods may be used for this purpose. Such numerical methods include different types of multivariate analysis. Computer programs for performing such methods are commercially available. They generally include a training which has to be performed by the manufacturer of the biosensor system using a plurality of samples with different hematocrit and measurements at different temperatures. The results of such training can then be programmed into the instrument for use during analysis of samples with unknown values of glucose and hematocrit concentrations on the basis of measurements at arbitrary temperatures within the temperature range for which the instrument was trained.
The temperature correction may in principle be based on only two values of the rising section of the I(t)-trace.
It is, however, preferred to include information regarding the curvature of the I(t)-trace into the correction algorithm. This requires measurement of at l.east three values of the analyte detection current at different points of time.
The rising section of the I(t).-trace shows a strong dependence on the measurement temperature and therefore values from the I(t)-trace in this section are valuable in particular for compensating temperature errors.
Generally the evaluation algorithm may additionally include measurement data generated at a point of time after the maximum of the I(t)-trace. Such data can in particular be of interest with respect to the sensitivity to changes of the analyte concentration and/or with respect to the quality of the hematocrit correction.
Preferably, however, the measurement of the I(t)-trace is terminated after the maximum of the I(t)-trace has been detected. This allows especially short measurement times if the temperature is relatively high (such as room temperature) and therefore the maximum is reached shortly after dose detect. On the other hand if a measurement has to be performed at a low temperature (e.g. during an out of house activity of the user) the instrument waits until the maximum has been detected thereby allowing a high accuracy measurement with a measurement time which is slightly larger but still as short as possible.
The biosensor system 1 shown in figure 8 consists of disposable analysis elements 2 and an evaluation instrument 3. Evaluation instrument 3 has a holder 4 for is receiving and holding an analysis element 2. When analysis element 2 is inserted into evaluation instrument 3, electrical connection is established by cooperation of electrical contacts 6, provided at analysis element 2, cooperating with mating contacts (not shown) of the instrument 3.
Analysis element 2 has a reaction zone 7 comprising a sample application opening 8 for applying a drop of blood sample. From here a narrow capillary space (not shown) extends to two electrodes 9,10,where the capillary space contains an analyte reactant (not shown). The dimensions of the capillary space are such that the sample liquid applied to the sample application opening 8 easily penetrates thereinto, disolves the analyte reactant and contacts both electrodes 9 and 10. Thereafter the above described reaction and evaluation steps take place.
In order to make the required electrical measurements and to derive the analyte concentration in the sample from the measured values instrument 3 contains the required electronic circuitry. This includes DC and AC-voltage sources, DC and AC current measurement circuits and a microprocessor-controlled evaluation system which controls the instrument operation and performs the steps of the evaluation algorithm by which the desired analyte concentration is derived from the measured electrical data.
Both with respect to the chemistry and design of the analysis element and with respect to the electronics of the instrument 3 system 1 is largely conventional (with the exception of particular features described herein).
Therefore no more detailed description is required.
Rather any further information which may be of interest can be taken from the appropriate literature as cited above.
Figure 9 is a schematic representation of the essential design features of an electronics which is suitable for simultaneously applying DC and AC-voltages to the electrodes of a biosensor and measuring the resulting DC
and AC current without mutual interference of the measurements.
The required DC and AC voltages are generated by respective voltage sources 20 and 21 and summed at a summing point 23 connected to the inverting input of an operational amplifier 24 which functions as a voltage follower. It provides at its output a low impedance source of the required DC and AC voltages which is connected via a shunt resistor 26 to one of the electrodes (namely electrode 10) of the analysis element 2. In order to allow monitoring of the voltage signal applied to electrode 10 a monitoring output 27 is provided.
Shunt resistor 26 belongs to an AC measurement circuit 25. The current flowing through shunt resistor 26 is measured by means of a differential amplifier 29 the output of which is proportional to the voltage drop at the shunt resistor 26 and thus of the current flowing therethrough. By means of an rms module 30 (rectifier-integrator) an output 31 corresponding to the rms-value of the AC-current flowing between the electrodes 9 and 10 is generated.
The DC-current flowing between electrodes 9 and 10 is measured by a separate current measurement circuit 32 which is independent of AC measurement circuit 25. It comprises a further operational amplifier 33 arranged and functioning as a current to voltage (I/V)-converter. Here the AC-signal is filtered out by a suitable filtering arrangement including (in the embodiment shown) two capacities 35,36 parallel to the input and parallel to the feedback resistor of the I/V-converter. Thus the DC
output 37 of the I/V-converter is proportional to the DC-current flowing between the electrodes 9 an.d 10.
All outputs 27,31 and 37 are digitized and transmitted to a microprocessor system of the,instrument for further processing as described.
Figure 10 shows measurements which were made with an electronics according to figure 9. The I(t)-traces shown refer to a plurality of different sample solutions. The corresponding DC-current scale is shown on the left ordinate. Simultaneously figure 10 shows impedance values (right ordinate) measured by an AC-measurement with switching of a plurality of different frequencies as shown. Again the same sample solutions were used and corresponding traces are marked by the same symbols. The figure clearly shows that the AC and DC-measurements are completely independent: There is not the slightest variation of the DC-signal at the points were the AC-5 frequency was switched and a corresponding sudden change of the AC-current was detected.
Figure 3 shows a typical shape of an I(t)-trace which results when the analysis voltage (DC-voltage required for the "reverse" electrode reaction) is applied not at the end of an incubation period required for the analyte specific forward reaction but rather at a much earlier point of time (preferably at "Dose Detect" or within at most 500 milliseconds, preferably at most 300 milliseconds and most preferably at most 100 milliseconds thereafter). Conveniently no special sample detect voltage is used but rather the, analysis voltage is applied before the sample is contacted to the electrodes.
The shape of the I(t)-trace shown in figure 3 is the result of a rather complicated combination of a plurality of influencing factors. In particular it depends on the speed of the enzymatic reaction which again depends not only on the concentration of the analyte but also on a plurality of potentially disturbing factors such as the temperature, and the diffusion of the reactants (which again is influenced by the temperature and the presence of particles, in particular blood cells, in the sample).
The product generated by the enzymatic reaction is simultaneously consumed by the electrode (reverse) reaction which is also influenced by a plurality of 5 factors, mainly by the concentration of the product of the enzymatic reaction in the vicinity of the electrode surface which again depends on the diffusion conditions in the sample liquid. In any case it is characteristic for the invention that the I(t)-traces include an initial 10 section in which the DC-current rises versus time. This section is designated rising section RS.
It has to be noted that an I(t)-trace measured in practice may have a more complicated shape, depending on the experimental setup. In particular it may not be , .monotonically rising in its section afterr application of the analysis voltage. Rather there may be an intermediate maximum as shown in dashed line in figure 3. Such a shape is, however, not due to the reaction sequence explained but rather due to transient behavior of the experimental setup, in particular of the measurement electronics. The rising section of the I(t)-trace as defined herein is the section preceding the maximum M which is due to the redox reaction sequence.
In the context of the invention it has been found that in spite of this complicated situation measurements of the analyte detection current which are made at a point of time within the rising section RS of the I(t)-trace (before it reaches its maximum M) can be used to improve the quality of the analytical result derived from the measurement. In particular by measuring the DC-current at at least two points of time in the rising section RS it is possible to compensate for errors caused by variations of the temperature of the reaction mixture. It is also il possible to derive the analyte concentration itself from current measurements in the rising section RS. Additional error corrections, in particular to compensate for errors caused by variations of the hematocrit values of the sample, are possible if also AC-measurements are performed during the rising section of the I(t)-trace.
Most preferably measurements of the DC-current and of the AC-current are performed simultaneously, i.e. there is an at least partial overlap of the measurement periods during which the DC- and AC-measurements are performed.
The invention achieves superior results regarding analytical tests at high speed and with very good analytical quality. In this context it has to be noted that generally enzymes providing a high specifity for a particular analyte react relatively slowly. With the classical method this increases the incubation period and therefore the time required for generating the analytical result. If enzymes are used which are optimized for short reaction times this generally requires a reduction of the specifity and thus*of the quality of the test. The invention allows to fulfill both requirements (high specifity and fast response) at the same time.
Figures 4 to 7 show results from experimental work performed in the context of the invention.
Figure 4 shows I(t)-traces as in figure 3, i.e. with the analysis voltage applied at dose detect, for a single glucose concentration (500mg/dl) but for varying values of the hematocrit and of the reaction liquid temperature.
The graph shows curves for three hematocrit values (20%, 45% and 70 %) and three temperatures (100, 25 and 40 C) each. The individual I(t)-traces are distinguished by using different symbols as shown in the column on the right side of the graphs. The same symbols are also used in figures 5 and 6. The data were derived from the original current measurements applying a baseline correction:
IBC (t) =I rg (t) -Cx (t-1/2x eo.1/t) wherein IBC . baseline corrected current Iorg . original Current C . empirical constant This baseline correction subtracts contributions caused by electrical charge carriers contained in the sample but not relating to the glucose concentration (t-1/2-term) and by the damping effect of the electronics (eo''/t-term) .
A further correction was necessary in view of the fact that it is not possible to prepare samples having exactly the desired values of the glucose concentration and of the hematocrit. Rather the samples had to be prepared by diluting available whole blood samples and adding glucose as required. Resulting minor deviations from the exact values shown in figure 4 were corrected by linear regression.
Figure 4 shows the following characteristics:
- Each of the curves includes a rising section essentially starting at dose detect (time 0) up to a maximum.
- The position of the maximum on the time scale depends on the temperature. The higher the temperature the earlier the maximum occurs (the drop of the I(t)-trace after the maximum is not visible for the 10 C-traces because it occurs after the maximum measurement time (5sec) shown). It is, however, possible to electronically detect a maximum with each of the traces).
- For each temperature the hematocrit causes a shift of the I(t)-trace: An increase of the hematocrit causes a decrease of the measured current values.
Figure 5 shows the same data as figure 4 but with a normalized time scale according to:
tnorm= t /tmax wherein tmax: time after dose detect at which the maximum of the respective 1(t) -trace occurs Figure 6 shows the results of AC-measurements with a frequency of 2 kHz versus time for the same samples as figure 4. The quantity shown is the admittance A
(reciprocal of impedance Z) measured in S(Ohm-1).. The curves show relatively strong initial variations which relate to the transient behavior of the measurement electronics. After at most two seconds, however, a stable value is achieved which shows pnly a very slow increase versus time which is due to a strong damping of the electronics used for the experiments.
Figure 7 shows the same data as figure 5 but after=a transformation of the current values according to:
Itf=lnorm x(1 + deltaZ x A) wherein deltaZ is calculated according to:
deltaZ= (Z-Zmed) wherein Z: impedance measured between the electrodes with an AC-voltage of 2 kHz at a point of time after the initial Z variations (e.g. at 3 sec in figure 6).
Zmed: Median of the Z values at this point of time A: Empirically determined weighing factor for optimum coincidence of the curves shown in figure 7.
The series of figures 4 to 7 shows that by relatively simple mathematical steps the complicated shape of the I(t)-traces shown in figure 4 can be reduced to essentially a single coinciding curve as shown in figure 7. In other words the I(t)-traces shown in figure 4 having a non-linear dependence from temperature and hematocrit can be transformed to I(t)-traces having a linear dependence on these disturbing factors by using a suitable transformation of the time axis.
These results show that - deviating from the prior art -reliable results, including correction for such disturbing factors as temperature and hematocrit, can be derived from amperometric measurements in which the enzymatic forward reaction takes place simultaneously with the reverse electrode reaction. It is not only possible but even favorable to use current measurement values from the rising section of the I(t)-traces. This again allows to substantially reduce the required time for performing a measurement because it is not necessary to introduce an incubation period for first performing the forward reaction and it is not even necessary to wait until the forward reaction and the reverse reaction lead to a steady state (in which the mediator is generated with the same speed in which it is consumed, i.e. its concentration is constant).
A simple way of performing an analysis of a sample having 5 an unknown glucose concentration comprises the following steps:
- An I(t)-trace as shown in figure 3 is measured.
- A baseline correction is performed.
- tma,, is determined.
10 - A current value measured at a time which corresponds to a predetermined point of time on the normalized time axis (e.g. 0.7 x tma,) is used for further evaluation.
- This current value is corrected on the basis of an AC
ls measurement.
- The resulting corrected current is transformed to a glucose concentration based on a calibration performed for the same point of time on the normalized time axis (e.g. 0.7 x tma,) This, however, is only one of several possibilities to evaluate measurements made in accordance with the invention. Figures 4 to 7 show,that the shape of the rising section of the I(t)-traces is characteristic for the combination of the influencing factors including a "variance" (the concentration of the analyte, e.g.
glucose) and a plurality of "covariances" (disturbing factors such as temperature and hematocrit). Therefore the effects of the covariances can be separated and consequently eliminated by measuring a plurality of values of the DC-current in the rising section of the I(t)-trace and by feeding these values into a suitable algorithm.
In place of the described mathematical-analytical method for example numerical methods may be used for this purpose. Such numerical methods include different types of multivariate analysis. Computer programs for performing such methods are commercially available. They generally include a training which has to be performed by the manufacturer of the biosensor system using a plurality of samples with different hematocrit and measurements at different temperatures. The results of such training can then be programmed into the instrument for use during analysis of samples with unknown values of glucose and hematocrit concentrations on the basis of measurements at arbitrary temperatures within the temperature range for which the instrument was trained.
The temperature correction may in principle be based on only two values of the rising section of the I(t)-trace.
It is, however, preferred to include information regarding the curvature of the I(t)-trace into the correction algorithm. This requires measurement of at l.east three values of the analyte detection current at different points of time.
The rising section of the I(t).-trace shows a strong dependence on the measurement temperature and therefore values from the I(t)-trace in this section are valuable in particular for compensating temperature errors.
Generally the evaluation algorithm may additionally include measurement data generated at a point of time after the maximum of the I(t)-trace. Such data can in particular be of interest with respect to the sensitivity to changes of the analyte concentration and/or with respect to the quality of the hematocrit correction.
Preferably, however, the measurement of the I(t)-trace is terminated after the maximum of the I(t)-trace has been detected. This allows especially short measurement times if the temperature is relatively high (such as room temperature) and therefore the maximum is reached shortly after dose detect. On the other hand if a measurement has to be performed at a low temperature (e.g. during an out of house activity of the user) the instrument waits until the maximum has been detected thereby allowing a high accuracy measurement with a measurement time which is slightly larger but still as short as possible.
The biosensor system 1 shown in figure 8 consists of disposable analysis elements 2 and an evaluation instrument 3. Evaluation instrument 3 has a holder 4 for is receiving and holding an analysis element 2. When analysis element 2 is inserted into evaluation instrument 3, electrical connection is established by cooperation of electrical contacts 6, provided at analysis element 2, cooperating with mating contacts (not shown) of the instrument 3.
Analysis element 2 has a reaction zone 7 comprising a sample application opening 8 for applying a drop of blood sample. From here a narrow capillary space (not shown) extends to two electrodes 9,10,where the capillary space contains an analyte reactant (not shown). The dimensions of the capillary space are such that the sample liquid applied to the sample application opening 8 easily penetrates thereinto, disolves the analyte reactant and contacts both electrodes 9 and 10. Thereafter the above described reaction and evaluation steps take place.
In order to make the required electrical measurements and to derive the analyte concentration in the sample from the measured values instrument 3 contains the required electronic circuitry. This includes DC and AC-voltage sources, DC and AC current measurement circuits and a microprocessor-controlled evaluation system which controls the instrument operation and performs the steps of the evaluation algorithm by which the desired analyte concentration is derived from the measured electrical data.
Both with respect to the chemistry and design of the analysis element and with respect to the electronics of the instrument 3 system 1 is largely conventional (with the exception of particular features described herein).
Therefore no more detailed description is required.
Rather any further information which may be of interest can be taken from the appropriate literature as cited above.
Figure 9 is a schematic representation of the essential design features of an electronics which is suitable for simultaneously applying DC and AC-voltages to the electrodes of a biosensor and measuring the resulting DC
and AC current without mutual interference of the measurements.
The required DC and AC voltages are generated by respective voltage sources 20 and 21 and summed at a summing point 23 connected to the inverting input of an operational amplifier 24 which functions as a voltage follower. It provides at its output a low impedance source of the required DC and AC voltages which is connected via a shunt resistor 26 to one of the electrodes (namely electrode 10) of the analysis element 2. In order to allow monitoring of the voltage signal applied to electrode 10 a monitoring output 27 is provided.
Shunt resistor 26 belongs to an AC measurement circuit 25. The current flowing through shunt resistor 26 is measured by means of a differential amplifier 29 the output of which is proportional to the voltage drop at the shunt resistor 26 and thus of the current flowing therethrough. By means of an rms module 30 (rectifier-integrator) an output 31 corresponding to the rms-value of the AC-current flowing between the electrodes 9 and 10 is generated.
The DC-current flowing between electrodes 9 and 10 is measured by a separate current measurement circuit 32 which is independent of AC measurement circuit 25. It comprises a further operational amplifier 33 arranged and functioning as a current to voltage (I/V)-converter. Here the AC-signal is filtered out by a suitable filtering arrangement including (in the embodiment shown) two capacities 35,36 parallel to the input and parallel to the feedback resistor of the I/V-converter. Thus the DC
output 37 of the I/V-converter is proportional to the DC-current flowing between the electrodes 9 an.d 10.
All outputs 27,31 and 37 are digitized and transmitted to a microprocessor system of the,instrument for further processing as described.
Figure 10 shows measurements which were made with an electronics according to figure 9. The I(t)-traces shown refer to a plurality of different sample solutions. The corresponding DC-current scale is shown on the left ordinate. Simultaneously figure 10 shows impedance values (right ordinate) measured by an AC-measurement with switching of a plurality of different frequencies as shown. Again the same sample solutions were used and corresponding traces are marked by the same symbols. The figure clearly shows that the AC and DC-measurements are completely independent: There is not the slightest variation of the DC-signal at the points were the AC-5 frequency was switched and a corresponding sudden change of the AC-current was detected.
Claims (18)
1. Biosensor system for electrochemical determination of an analyte concentration in a blood sample, the system comprising:
an analyte reactant and at least two electrodes, wherein the system is adapted for:
mixing of the analyte reactant with the blood sample to form a reaction liquid and contacting of the reaction liquid to the electrodes for measurement of an analyte detection current flowing between the electrodes in response to a DC-voltage applied thereto, the analyte detection current being characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction taking place at an electrode of said at least two electrodes including transfer of electrons through an electrode surface; and electronic circuitry comprising a DC-voltage source for applying a required DC-voltage to the electrodes and a measurement and evaluation electronics for measuring a plurality of values of a curve of the analyte detection current versus time over a plurality of measurement times and for deriving the analyte concentration therefrom by means of an evaluation algorithm characterized in that the electronic circuitry is adapted for:
application of the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously in accordance with a relationship defined by an analyte detection current versus time curve which comprises a rising section in which the analyte detection current rises versus time, measuring of the analyte detection current at at least two points of time in the rising section of the curve; and using the values of the analyte detection current resulting from said at least two measurements in the evaluation algorithm for a temperature error compensation.
an analyte reactant and at least two electrodes, wherein the system is adapted for:
mixing of the analyte reactant with the blood sample to form a reaction liquid and contacting of the reaction liquid to the electrodes for measurement of an analyte detection current flowing between the electrodes in response to a DC-voltage applied thereto, the analyte detection current being characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction taking place at an electrode of said at least two electrodes including transfer of electrons through an electrode surface; and electronic circuitry comprising a DC-voltage source for applying a required DC-voltage to the electrodes and a measurement and evaluation electronics for measuring a plurality of values of a curve of the analyte detection current versus time over a plurality of measurement times and for deriving the analyte concentration therefrom by means of an evaluation algorithm characterized in that the electronic circuitry is adapted for:
application of the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously in accordance with a relationship defined by an analyte detection current versus time curve which comprises a rising section in which the analyte detection current rises versus time, measuring of the analyte detection current at at least two points of time in the rising section of the curve; and using the values of the analyte detection current resulting from said at least two measurements in the evaluation algorithm for a temperature error compensation.
2. System according to claim 1, characterized in that it comprises:
a disposable analysis element (2) including the analyte reactant and the at least two electrodes (9,10) and an evaluation instrument (3) adapted for evaluation of the analysis element (2) and having a holder (4) for receiving the analysis element (2) and electrical contacts establishing an electrical connection to the analysis element (2) inserted into the holder (4) thereby connecting the electrodes (9,10) of the analysis element (2) to the electronics circuitry.
a disposable analysis element (2) including the analyte reactant and the at least two electrodes (9,10) and an evaluation instrument (3) adapted for evaluation of the analysis element (2) and having a holder (4) for receiving the analysis element (2) and electrical contacts establishing an electrical connection to the analysis element (2) inserted into the holder (4) thereby connecting the electrodes (9,10) of the analysis element (2) to the electronics circuitry.
3. Biosensor system according to claim 2, characterized in that the analysis element is a capillary type element including a space of capillary dimensions comprising the electrodes (9,10) and the analyte reactant.
4. System according to claim 1, characterized in that the electronic circuitry is adapted for measuring the analyte detection current at at least three points of time in the rising section to generate information about a shape of the rising section and for using the at least three values of the analyte detection current resulting from said at least three measurements in the evaluation algorithm for the temperature error compensation.
5. System according to claim 1 characterized in that the electronic circuitry is further adapted for applying an AC-voltage to the electrodes and for measuring a resulting AC-current for error compensation.
6. System according to claim 5, characterized in that the electronic circuitry is further adapted for performing the measuring of the AC-current simultaneously with a measuring of the analyte detection current in the rising section.
7. System according to claim 6, characterized in that the sample is whole blood and the error compensation based on a resulting AC-current measurement comprises a compensation for variations of a hematocrit value of the whole blood sample.
8. System according to claim 1, characterized in that the evaluation algorithm comprises determination of a maximum of the analyte detection current versus time curve and use of a value of the analyte detection current versus time curve which has been measured at a point of time corresponding to a predetermined ratio to the point of time at which the maximum occurs.
9. System according to claim 6, characterized in that the electronic circuitry comprises two independent current measurement circuits (25,32) for measuring the AC-current and the DC-current, respectively.
10. System according to claim 9, characterized in that one of the current measurement circuits (25) is located in a first connecting line leading to one of said at least two electrodes and the other current measurement circuit (32) is located in a second connecting line leading to the other of said at least two electrodes.
11. System according to claim 9, characterized in that the AC-current measurement circuit (25) comprises a shunt resistor coupled to a differential amplifier.
12. System according to claim 9, characterized in that the DC-current measurement circuit (32) comprises an operational amplifier based voltage to current converter.
13. Method for electrochemical determination of an analyte concentration in a blood sample, comprising the steps of:
mixing an analyte reactant with the sample to form a reaction liquid, contacting the reaction liquid to a pair of electrodes, measuring a plurality of values of a curve of an analyte detection current versus time over a plurality of measurement times, the analyte detection current flowing between the electrodes in response to a DC-voltage applied thereto and being characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction taking place at an electrode of said pair including transfer of electrons through an electrode surface; and deriving the analyte concentration by means of an evaluation algorithm comprising use of said plurality of values of the analyte detection current, characterized in that:
the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously resulting in the analyte detection current versus time curve which comprises a rising section in which the current rises versus time; and the analyte detection current is measured at at least two points of time in the rising section of the curve and the values of the analyte detection current resulting from said at least two measurements are used in the evaluation algorithm for temperature error compensation.
mixing an analyte reactant with the sample to form a reaction liquid, contacting the reaction liquid to a pair of electrodes, measuring a plurality of values of a curve of an analyte detection current versus time over a plurality of measurement times, the analyte detection current flowing between the electrodes in response to a DC-voltage applied thereto and being characteristic of the analyte concentration in the sample by virtue of a reaction sequence comprising an analyte specific reaction taking place in the reaction liquid and an electrode reaction taking place at an electrode of said pair including transfer of electrons through an electrode surface; and deriving the analyte concentration by means of an evaluation algorithm comprising use of said plurality of values of the analyte detection current, characterized in that:
the DC-voltage is applied to the electrodes during the analyte-specific reaction whereby the analyte-specific reaction and the electrode reaction take place simultaneously resulting in the analyte detection current versus time curve which comprises a rising section in which the current rises versus time; and the analyte detection current is measured at at least two points of time in the rising section of the curve and the values of the analyte detection current resulting from said at least two measurements are used in the evaluation algorithm for temperature error compensation.
14. Method according to claim 13, characterized in that the analyte detection current is measured at at least three points of time in the rising section to generate information about a shape of the rising section and the at least three values of the analyte detection current resulting from said at least three measurements are used in the evaluation algorithm for the temperature error compensation.
15. Method according to claim 13, characterized in that an AC-voltage is applied to the electrodes and the resulting AC-current is measured for error compensation.
16. Method according to claim 15, characterized in that a measurement of the AC-current is performed simultaneously with a measurement of the analyte detection current in the rising section.
17. Method according to claim 16, characterized in that the sample is whole blood and the error compensation based on a resulting AC-current measurement comprises a compensation for variations of a hematocrit value of the whole blood sample.
18. Method according to claim 13, characterized in that the evaluation algorithm comprises determination of a maximum of the analyte detection current versus time curve and use of a value of the analyte detection current versus time curve which has been measured at a point of time corresponding to a predetermined ratio to the point of time at which the maximum occurs.
Applications Claiming Priority (3)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| EP03008143.4 | 2003-04-08 | ||
| EP03008143A EP1467206A1 (en) | 2003-04-08 | 2003-04-08 | Biosensor system |
| PCT/EP2004/002646 WO2004090533A1 (en) | 2003-04-08 | 2004-03-13 | Biosensor system |
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| CA2521370A1 CA2521370A1 (en) | 2004-10-21 |
| CA2521370C true CA2521370C (en) | 2010-07-27 |
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| Application Number | Title | Priority Date | Filing Date |
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| CA2521370A Expired - Fee Related CA2521370C (en) | 2003-04-08 | 2004-03-13 | Biosensor system |
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|---|---|
| US (1) | US7638033B2 (en) |
| EP (2) | EP1467206A1 (en) |
| JP (1) | JP4381413B2 (en) |
| AT (1) | ATE393911T1 (en) |
| CA (1) | CA2521370C (en) |
| DE (1) | DE602004013438T2 (en) |
| ES (1) | ES2304608T3 (en) |
| WO (1) | WO2004090533A1 (en) |
Cited By (1)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US8721851B2 (en) | 2008-11-28 | 2014-05-13 | Panasonic Healthcare Co., Ltd. | Sensor chip, biosensor system, method for measuring temperature of biological sample, method for measuring temperature of blood sample, and method for measuring concentration of analyte in blood sample |
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- 2004-03-13 CA CA2521370A patent/CA2521370C/en not_active Expired - Fee Related
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- 2004-03-13 WO PCT/EP2004/002646 patent/WO2004090533A1/en active IP Right Grant
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Cited By (2)
| Publication number | Priority date | Publication date | Assignee | Title |
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| US8721851B2 (en) | 2008-11-28 | 2014-05-13 | Panasonic Healthcare Co., Ltd. | Sensor chip, biosensor system, method for measuring temperature of biological sample, method for measuring temperature of blood sample, and method for measuring concentration of analyte in blood sample |
| US9658182B2 (en) | 2008-11-28 | 2017-05-23 | Panasonic Healthcare Holdings Co., Ltd. | Method for measuring concentration of analyte in blood sample, and biosensor system |
Also Published As
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| JP2006522923A (en) | 2006-10-05 |
| EP1613955B1 (en) | 2008-04-30 |
| ATE393911T1 (en) | 2008-05-15 |
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| US7638033B2 (en) | 2009-12-29 |
| CA2521370A1 (en) | 2004-10-21 |
| DE602004013438D1 (en) | 2008-06-12 |
| ES2304608T3 (en) | 2008-10-16 |
| DE602004013438T2 (en) | 2009-04-23 |
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| WO2004090533A1 (en) | 2004-10-21 |
| EP1613955A1 (en) | 2006-01-11 |
| EP1467206A1 (en) | 2004-10-13 |
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