Method and sensor for determining d-lactate

Abstract

A method and a sensor for determining an analyte concentration in a sample are described. The sensor comprises at least two measuring chambers, which can be filled with a sample via a sample-receiving region of the sensor, wherein the at least two measuring chambers are formed on a support and comprise a first three-electrode arrangement, a four-electrode conductivity arrangement and a second three-electrode arrangement. The measuring signal is corrected with the aid of these electrode arrangements in order to minimize influences of electrochemically active substances, and other contaminating substances, in the sample. The system is suitable in particular for determining the concentration of D-lactate in clinical sample material, for example synovial fluid, cerebrospinal fluid, urine or blood.

Description

[0001] Method and sensor for the determination of D-lactate

[0002] The invention relates to a method and a sensor for determining an analyte concentration in a sample, in particular D-lactate.

[0003] Technological background

[0004] Whole blood measurements of analytes such as lactate or glucose using enzymatic-electrochemical sensors, especially disposable sensors, are influenced by a number of parameters resulting from both ambient conditions and the sample matrix itself. Major causes of erroneous measurements can include contamination during sample collection (e.g., contamination of a synovial fluid sample by blood), the sample's hematocrit, redox-active endogenous metabolites and medications, and, above all, the ambient temperature. Temperature influences not only the enzymatic indication reaction, but also the sample's viscosity, the dissolution behavior of the reagent layer, the diffusivity of the substances involved in the indication reaction, including the diffusion-inhibiting effect of the hematocrit, and the reference potential.

[0005] Lactate is a product of anaerobic glycolysis formed from pyruvate by the action of lactate dehydrogenase, or LDH for short. LDH enzymes are typically specific for a particular stereoisomer of lactate. Since D-specific LDH does not occur in humans but is present in most bacteria, D-lactate potentially represents a specific biomarker for bacterial growth that can accumulate in the body up to millimolar amounts. D-lactate is produced only in relatively small amounts in humans via the methylglyoxal pathway and can only be metabolized relatively inefficiently to pyruvate by the mitochondrial enzyme D-2-hydroxyacid dehydrogenase; the resulting serum concentrations are typically in the nano- to micromolar range.While L-lactate is a widely established metabolic biomarker frequently tested in diagnostic laboratories as an essential part of routine clinical chemistry workup, D-lactate is less frequently tested, and most diagnostic lactic acid tests in the clinical laboratory are specific for the L-isomer. Nevertheless, D-lactate is known to be a valuable biomarker because it is a relatively specific indicator of bacterial growth and has been clinically associated with a number of different diseases. The presence of D-lactate in the blood is usually indicative of bacterial overgrowth in the gastrointestinal tract as a result of a malabsorption disorder, such as short bowel syndrome. The resulting D-lactic acidosis can lead to mental disturbances ranging from drowsiness to coma, encephalopathy, or other neurological symptoms.Even healthy children suffering from gastroenteritis can develop D-lactic acidosis. Abnormally high D-lactate concentrations are considered a sign of sepsis, ischemia, or trauma.

[0006] When designing disposable sensors with temperature correction, a calibration curve is typically recorded with spiked samples at a standard temperature, such as 25°C, which later forms the basis for each measurement. In parallel, a set of calibration curves is repeated at temperatures expected in the subsequent measurement environment. This temperature range is typically between 5°C and 40°C. An ambient temperature lower than the nominal temperature results in a lower signal current at a lower base current value, while an ambient temperature higher than the nominal temperature results in a higher signal current at a higher base current value.The resulting base current and slope values ​​of the calibration curves, which were then obtained for selected temperature values ​​within this range, are used, using the standard calibration curve parameters, to develop an algorithm that can be used to correct the current concentration value for the temperature-related deviation. However, for this purpose, it is advisable that the current ambient temperature is known or can be measured as accurately as possible and close to the sensor.

[0007] The most common technical solution for compensating for temperature influences involves the use of a miniaturized PT100 temperature sensor, which is located in the handheld measuring device in close proximity to the plugged-in disposable sensor. Its calibrated temperature measurement is used via an algorithm to correct the measured value. Since the measuring device, unlike the disposable sensor, has a relatively high heat capacity, significant differences can occur between the measured temperature on the device and the temperature in the sensor's measuring chamber, particularly during rapid changes in ambient temperature, due to heat generation from components, or after charging an internal battery.Furthermore, depending on the time between lancing and application of the sample to the sensor, if the ambient temperature deviates significantly from the standard measurement temperature, the blood droplet may result in a measurement temperature on the sensor that differs significantly from the temperature recorded on the measuring device. In both cases, incorrect temperature compensation of the measured value will occur.

[0008] Other known solutions from the prior art describe the use of an amperometric two-electrode arrangement consisting of a working (anode) and a reference (cathode) electrode, which is mounted separately next to the amperometric measuring system and is not covered with the detection reagent. After applying an extremely high positive polarization voltage, ranging from 1 to 3 V, a signal current is generated whose magnitude is independent of the Hct and analyte and essentially depends only on the sample or ambient temperature. The diffusion of reagent to the "temperature electrodes" after filling the measuring chamber can prove problematic.

[0009] Furthermore, it is known to establish a defined relationship between the ambient temperature, which is determined via a physical characteristic of the sample by means of additionally arranged electrodes, and, if necessary, the temperature measured by the measuring device in order to achieve compensation for the temperature influence.

[0010] It is also known that, by using a temperature sensor in the instrument but in combination with a series of heat-generating pulses of defined duration and amplitude, which lead to a defined temperature increase via metallic conductors on the disposable sensor, the probable temperature at the sensor's indication area can be determined from the difference. In other cases, a temperature sensor is positioned as part of the sensor contact on the underside of the sensor substrate. Its value is recorded during the test measurement and serves to correct the temperature-dependent measured value deviation.

[0011] A technical solution for a disposable sensor for voltammetric analyte concentration measurement utilizes admittance and phase angle measurements, in which up to four different frequencies up to 20 kHz are applied briefly. The admittance is hematocrit and temperature dependent, but only a hematocrit dependence was found for the phase angles. Using different frequencies to increase reliability, an algorithm can be developed from the empirical determination of the dependencies between the two values. This algorithm determines the reaction zone temperature in addition to the hematocrit, so that both variables can be used to correct the enzymatic-electrochemical analyte determination.

[0012] Furthermore, a temperature-dependent measurement signal can be determined by applying the polarization voltage, first reducing it below a threshold required to maintain the electrochemical indication reaction, and measuring an offset current. This is followed by a further reduction of the polarization voltage to an even lower level, so that a second offset current is measured. The difference between the two offset currents exhibits a temperature dependence, so that with appropriate calibration, a temperature value is obtained that is used to compensate for the temperature-dependent indication reaction. However, it cannot be ruled out that the hematocrit of the sample causes an additional dependence.

[0013] Another cause of erroneous measurements can be contamination of clinical sample materials by blood components (erythrocytes) or, in the case of whole blood samples, by the hematocrit of the blood sample. Contamination of the cerebrospinal fluid (CSF) or synovial fluid with increasing blood volume can lead to falsely elevated protein or metabolite concentrations. Therefore, it may be advantageous to consider the amount of blood contamination when interpreting the results of concentration determinations in CSF or synovial fluid. The hematocrit (Hct) represents the volume fraction of red blood cells in the blood, which accounts for approximately 99% of the cellular components in whole blood. In healthy adults, the hematocrit is between 40% and 48% (men) and 36% to 42% (women).However, depending on the genetic predisposition of a subject, their age, gender, state of health and physical activity, values ​​between 20% and 70% can also occur.

[0014] Erythrocytes are oval-shaped cells with a diameter of 2 to 30 μm. They have a significantly lower water content (74.6 g / dL, w / v) than plasma (94.2 g / dL, w / v). For example, when measuring glucose, which can easily penetrate the erythrocyte membrane, in whole blood with a normal Hct, this leads to a reading approximately 10% lower than that measured in plasma.

[0015] If the metabolites are transported into the erythrocytes via transporter proteins, resulting in a difference between their concentration in plasma and in the erythrocytes, the deviation can be significantly greater. For example, for L-lactate, the concentration in the erythrocytes is described to be between 50% and 70% of the plasma value.

[0016] When designing single-use sensors, the measured value for the normal hematocrit range is calibrated to the whole blood value so that Hct-related measurement errors lie within a defined, previously determined concentration range and are stated in the specification.

[0017] However, for hematocrit values ​​outside the normal range, this measurement error caused by the red blood cell volume can increase considerably.

[0018] In addition, the Hct contributes to further error-causing effects in voltammetric measurements of disposable sensors that contain an oxidoreductase for the specific reaction with the target analyte and a redox mediator for electron transfer between the oxidoreductase and the working electrode surface: (i) Due to the lower conductivity of erythrocytes compared to plasma, stationary measurement in an electrochemical measuring cell leads to an increase in the measuring cell resistance and thus to a lower current. This effect can be largely compensated for by using a potentiostatic three-electrode arrangement, (ii) Due to the "particle character" of the erythrocytes, the diffusion processes are hindered, which impairs both the enzymatic substrate and theThis applies to both the analyte and the electron mediator, which transports the electrons for or from the enzymatic redox reaction between the enzyme and the electrode surface. The same applies to the corresponding mass transport at the reference or counter electrode. Due to the diffusion barrier, a lower Faraday current is generated compared to plasma measurement or compared to the measurement of a sample not contaminated by erythrocytes.

[0019] As a result, the measured analyte concentration will be too low for samples with high hematocrit and too high for samples with low hematocrit.

[0020] To compensate for the hematocrit value, it is known that two electrodes of the amperometric measuring chain are subjected to an alternating current component in parallel to the direct voltage to enable an impedance measurement. From the admittance and / or the phase angle, which are determined at different frequencies if necessary, a factor for the hematocrit dependence is determined and used to correct the amperometric measurement. It is also known to perform two impedance measurements in the sample measuring chamber of a disposable sensor, the frequencies of which can differ by a factor of two to 100, with the second frequency being greater than 20 kHz. The impedance measurement is performed using a two-electrode arrangement located near the sample intake opening, so that a first impedance measurement is taken immediately upon sample entry and before the reagent mixture dissolves, and a second impedance measurement is taken after the reagent mixture dissolves.A multiple regression analysis is carried out from the two measured values ​​using empirically determined functions, which also depend on cell parameters and the reagent system, and the hematocrit value is calculated, which serves to compensate for the analyte-dependent useful signal.

[0021] Furthermore, the state of the art includes the design of a measuring chamber with two measuring zones, each with two consecutively arranged electrode pairs, the first two of which are covered with an analyte-detecting reagent layer, forming an amperometric two-electrode arrangement. This measuring zone is followed by a separating element in the form of a polymer layer, which is then followed by the second measuring zone with the second pair of electrodes. The latter are operated with an alternating voltage and serve to measure the conductivity of the sample. The hematocrit-dependent conductivity measurement is used to compensate for the hematocrit-dependent amperometric concentration measurement.

[0022] Furthermore, a sensor arrangement with a capillary gap measuring chamber is known, in which two lateral spacer walls consist of a conductive material, to which an alternating current is applied in order to determine the conductivity of the sample and thus the hematocrit value, which serves to correct the analyte-dependent amperometric measurement. It is also known to provide a capillary measuring chamber with three consecutively arranged identical circular electrodes, which represent two working electrodes and the reference electrode, and are microfluidically connected to one another via a gap-like channel arranged above them. One of the working electrodes and the reference electrode contain an identical reagent mixture, so that after the sample is drawn through the connecting channel between the two electrodes, the hematocrit-dependent conductivity of the sample is measured, which is used to correct the amperometric measurement.

[0023] When measuring the conductivity of the sample, however, it must also be taken into account that analytes such as L- or D-lactic acid or the resulting L-lactate or D-lactate are present in dissociated form and cause an increase in conductivity with increasing concentration, so that the Hct determined via conductivity measurement may be erroneous without taking the analyte concentration into account.

[0024] Other known solutions utilize pulsed potential and potential scan methods, special electrode arrangements in a measuring cell, and additional redox indicators to quantify diffusion-dependent influences in voltammetric measurements in blood samples, such as those emanating from hematocrit. For example, diffusion-determining scan phases during a cyclic scan can be used to determine the hematocrit fraction, for example, by determining the ratio of peak current to plateau current, which is analyte-independent and essentially depends only on the component of the sample that determines diffusion.

[0025] Furthermore, using differential pulse voltammetry (DPV) for blood concentration determination, the resulting current curves can be evaluated even below the diffusion-limiting peak current, largely eliminating the diffusion influence of the hematocrit on the measurement result. Short, high-frequency pulses are applied to maintain the diffusion layer of the reagent layer.

[0026] Another well-known technique is square-wave voltammetry (SWV), which uses an additional redox mediator to detect a redox-mediated enzymatic detection reaction in blood samples. Its formal redox potential is far enough from that of the indication mediator that the mediator cannot react with either the indication mediator or the oxidoreductase. This redox mediator acts as an internal standard, which, like the signal-generating system, is influenced by hematocrit and temperature via the diffusion rate. However, it is detected independently of the substrate and can thus provide hematocrit information.

[0027] The mobility or diffusivity of the redox mediator, which is significantly influenced by the Hct, can be determined as follows: After the voltammetric-enzymatic measurement of the analyte concentration, the polarization voltage is switched off, and the voltage drop, which is essentially determined by the mobility of the redox mediator, is monitored potentiometrically between the working and reference electrodes. The rate of the drop is thus a parameter for the indirect determination of the hematocrit value, which serves to correct the concentration measurement and, if necessary, also allows a distinction between control solution and sample. The mobility of the redox mediator can be determined by applying the redox mediator to the reference electrode and measuring the resulting reduction current after applying a voltage to the working electrode that is negative compared to the reference electrode.The measuring current depends exclusively on how quickly the redox mediator can diffuse to the working electrode. The lower the current, the greater the diffusion barrier or the higher the proportion of erythrocytes in the sample, resulting in a hematocrit-dependent signal that can be used to correct the actual concentration-dependent measurement signal.

[0028] Based on the hematocrit-related viscosity change of the sample, which can lead to correspondingly different filling times of the measuring chamber of a capillary gap sensor, a hematocrit-dependent signal is to be determined, which is particularly prone to errors in the case of fluctuating ambient temperatures.

[0029] Other solutions, in addition to the use of the electron mediator required for the enzymatic redox reaction, describe the use of a second redox mediator that does not react with the first and differs in its formal redox potential by at least 200 mV from the first redox mediator. The amperometric detection of the reduction or oxidation of this mediator will then depend significantly on its diffusivity in the measuring cell, which is determined by the hematocrit concentration. The resulting hematocrit-dependent current signal, which is quantified using chronoamperometry, SWV, DPV, or CV (cyclic voltammetry) and Hct-calibrated according to an empirically determined algorithm, serves to compensate for the hematocrit dependence of the analyte-dependent signal. Another technical solution assumes that the charging current immediately after application of the polarization voltage is largely determined by the hematocrit.If this value is related to the subsequently occurring Faraday current by forming a quotient, a hematocrit-dependent value is also obtained.

[0030] Other amperometric measurement procedures, which are known from technical solutions, evaluate transition states after the application of the polarization voltage, which is changed in amplitude and sign, in order to determine the Hct component in addition to the analyte-dependent signal.

[0031] In another method, two polarization voltages of opposite sign are applied successively over a period of 1 to 20 s, with the first being negative and the second being positive. From the resulting "transient currents," which depend on the analyte concentration and the electrochemical cell constant, and thus on the Hct, a hematocrit-corrected analyte concentration is determined using a three-dimensional calibration graph. Specifically, the graph is based on the ratio of negative and positive current values ​​and analyte concentration curves calibrated against a reference system, each recorded at different hematocrit values.The hematocrit-dependent information is derived from the fact that immediately after applying a suitable polarization voltage, the resulting current pulse initially consists primarily of charging current and a Faradaic current component. Toward the end of the pulse, the Faradaic component predominates, which is characterized by the Cotrell equation. Since the effective diffusion coefficient of the analyte or a mediator involved in the indication reaction is influenced by the hematocrit, a hematocrit-dependent value can be determined from a suitable ratio of current segments of the initial signal current at a specific polarization voltage or two consecutively applied polarization voltages with opposite signs. This value can be used to correct the desired signal.However, since the hematocrit influence varies depending on the magnitude of the anodically generated current of the analyte, the analyte concentration must also be taken into account when calculating the hematocrit correction. This fact, which must be considered for every electrochemical cell constant, results in the need for a three-dimensional set of curves.

[0032] To obtain a hematocrit-independent analyte measurement using an amperometric two-electrode arrangement consisting of a microelectrode array as the working electrode and a reference electrode, the polarization voltage is applied at intervals of up to three times, each lasting between 0.1 s and 1 s, with recovery phases in between. The polarization voltage is applied at near 0 mV for a duration of 0.05 s and 1 s. The resulting chronoamperometric current curve section is recorded, with the logarithmic current values ​​at the end of each interval being used for evaluation by plotting them against the logarithm of the measurement times. It was found that the resulting intersection with the ordinate corresponds to a virtually hematocrit-independent useful signal.

[0033] Another approach using this amperometric electrode arrangement aims to measure a so-called test current value immediately after the polarization voltage is applied and a steady-state current curve is reached. Plotting the respective quotients against the inverse square root of time yields a linear slope that corresponds to the diffusion coefficient and is used for the hematocrit-compensated calculation of the analyte value.

[0034] A more general approach, which also utilizes the effects described above, is based on an amperometric two-electrode arrangement consisting of an anode and a cathode, which are arranged plane-parallel to each other, with the resulting gap representing the measuring chamber. Polarization voltages of opposite signs are applied in two consecutive intervals. The level of the voltage, referred to as the test potential, which lies between -100 mV and -600 mV or +100 mV and +600 mV, is intended to enable partial reduction or oxidation at the electrode surfaces. During the first interval, a voltage is applied which causes partial oxidation of the reduced mediator at the second electrode, i.e. the first electrode is cathodically active and reduces the mediator. During the second interval, partial oxidation of the mediator, which may have been reduced by the enzymatic indication reaction, occurs at the first electrode.Accordingly, for example, a negative and then a positive potential is applied to the first electrode. A preliminary measured value for the analyte concentration is first determined from the signal current resulting from one of the two intervals. The hematocrit-dependent error source is then identified, and the analyte value is corrected, taking into account previously determined factors such as different concentration ranges for analyte and hematocrit values ​​and evaluation criteria for high or low analyte concentration at high or low hematocrit values, which are based on different empirically determined functions.

[0035] In other solutions, two consecutive polarization voltages of the same polarity are applied to an amperometric two-electrode arrangement. After triggering, a low anodic voltage between 10 mV and 150 mV is applied within the first second, followed by a higher anodic voltage between 250 mV and 350 mV. The resulting current pulse 450 ms to 530 ms after triggering is evaluated to determine the hematocrit, and the subsequent pulse with a higher polarization voltage is used to determine a preliminary analyte concentration value.

[0036] Hematocrit is determined using a set of curves previously determined with defined analyte and hematocrit concentrations. Different, empirically determined equations were determined for three pre-stored concentration ranges, each covering the entire analytically relevant range of the analyte, which are used to correct the hematocrit of the measured analyte concentration.

[0037] A disadvantage of the methods described above, however, is that they are based on complex calculations using empirically determined three-dimensional sets of curves, which have to take into account the diffusion-dependent Faraday current component of the signal current or the level of the analyte concentration.

[0038] To prevent the influence of electrochemically active substances from metabolic reactions or medications on analyte measurement, a number of technical solutions are known that are intended to minimize, compensate for, or prevent their influence. Practical solutions describe additional protective layers, semipermeable or porous membrane layers, applied over the reaction layer. These layers are intended to prevent or at least significantly reduce the diffusion of the interfering substances to the electrode system during the comparatively short measurement time. However, separate detection of the interfering substances is more effective, especially for patients in clinical settings, since little-known electrochemically active substances from medication intake or cumulative effects can be expected here, for which the diffusion barrier layers are not designed.

[0039] In one series of disclosures, an additional working electrode is used, which is operated at an overpotential, so that the interfering substances are oxidized at the electrode to a significantly greater extent than the reduced mediator. The factor, determined from the measurement signals of the two working electrodes using an empirical equation, serves to correct the analyte measurement value. However, this approach does not take into account that electrochemically active substances are not only oxidized directly at the surface of the working electrode, but also react with the redox mediator.

[0040] One of the first commercial disposable sensor devices based on an electrochemical-enzymatic indication reaction uses a tissue layer structure with hydrophilic and hydrophobic tissue properties over an electrode array. A sample collection port is provided in the upper layer. The first tissue layer is additionally impregnated with erythrocyte-aggregating polymeric components, intended to ensure effective Hct retention. An additional working electrode, coated with an enzyme-free reagent, serves to measure electrochemically active, interfering components. This variant still required a comparatively large sample volume and, due to the time-dependent aggregation, a long measurement time of 20 s.

[0041] A known technical solution for sensor disposables for the determination of glucose, lactate and creatinine uses an additional working electrode in a capillary gap arrangement, which serves exclusively to detect interference currents from the oxidation of ascorbate, uric acid, bilirubin, paracetamol or other electrochemically active substances in the blood sample, so that the total signal current can be corrected by this amount.

[0042] It is also known to use an overvoltage in the form of an exciting waveform at the working electrode to generate nonlinear current signals. Resolving the signal distribution using a vector projection method yields a multitude of real and imaginary components, multiple Fourier coefficients, or multiple frequencies, which are calibrated to a reference current signal of an analyte or interfering components, allowing either the analyte itself or the interfering substances to be selectively determined.

[0043] It is also known that, to detect electrochemically active substances, transient currents are measured immediately after a change in the polarization voltage, which is performed immediately after filling the measuring chamber. Up to three different polarization voltages of opposite polarity are applied to the working electrode in quick succession but at defined times. After adjusting the respective transient currents, they are compared, and the proportion of interference currents caused by electrochemically active substances is subtracted.

[0044] A disadvantage of all previously known technical solutions is the fact that indirect measurement methods are used to correct for interference with the analyte signal, which are often associated with large measurement value scatter. Correcting the considerable influence of temperature on hematocrit determination, as well as the concentration of ionically conductive components in the sample, such as a dissociated analyte or dissociated endogenous or exogenous metabolites on hematocrit determination, is not described in existing technical solutions. Furthermore, only an external temperature sensor is used for temperature measurement, which is integrated into a connected measuring device, and can thus lead to incorrect temperature corrections.

[0045] An object of the invention is therefore to overcome the aforementioned disadvantages and to provide a sensor and a method for the highly accurate determination of D-lactate concentration in various clinical sample materials, for example synovial fluid, cerebrospinal fluid, urine, or blood, in order to enable a more accurate enzymatic-electrochemical analyte determination during the measurement procedure, compensating for the actual temperature, the conductivity of the sample, and the electrochemically active substances contained therein, so that an adequate correction of these interferences on the plasma-related analyte measurement value to be determined can be achieved. Further solved problems will become apparent from the following description. The object of the invention is achieved according to the independent claims. The subclaims represent preferred embodiments.

[0046] Summary of the invention

[0047] A first aspect of the invention relates to a method for determining an analyte concentration, in particular of D-lactate, in a sample, in particular in clinical sample material, for example synovial fluid, cerebrospinal fluid, urine, or blood, using a sensor. In one step, at least two measuring chambers are filled with a sample via a sample receiving area of ​​the sensor, wherein the at least two measuring chambers are formed on a carrier of the sensor and at least two measuring chambers comprise a first three-electrode arrangement, a four-electrode conductivity arrangement, and a second three-electrode arrangement. In an optional step, an ambient temperature is measured after filling the measuring chambers using a temperature measuring resistor applied to the carrier. The method further comprises the step of determining an ionic conductivity of the sample using the four-electrode conductivity arrangement.The method further comprises the step of determining an interference charge of electrochemically active substances in the sample using the first three-electrode arrangement. The method further comprises the step of enzymatically electrochemically determining an analyte charge in the sample using the second three-electrode arrangement. The method further comprises determining an analyte concentration using pre-stored calibration curves, optionally the determined ambient temperature, and the determined analyte charge. The method further comprises determining an interference concentration using pre-stored calibration curves, optionally the determined ambient temperature, and the determined interference charge; and further determining a contamination value using pre-stored calibration curves, optionally the determined ambient temperature, and the determined ionic conductivity.The method further comprises correcting the (optionally temperature-corrected) analyte concentration and the (optionally temperature-corrected) interference concentration to a contamination-corrected analyte concentration and a contamination-corrected interference concentration using pre-stored calibration curves and the previously determined (optionally temperature-corrected) contamination value. The method further comprises the step of determining the interference-corrected analyte concentration by subtracting the interference concentration from the analyte concentration.

[0048] The analyte can preferably be D-lactate. Electrochemically active substances can be, for example, uric acid, bilirubin, ascorbate, or paracetamol. Interference-corrected means that the analyte concentration is based on the value of a corresponding sample without contaminating substances. Such contaminating substances can be, for example, erythrocytes and their components, also referred to herein as "hematocrit," or leukocytes and their components.

[0049] The method is preferably carried out in vitro, i.e. outside the human or animal body.

[0050] Potentiostatic three-electrode arrangements are preferably used as indication systems for both measurements. The measurements can preferably be carried out voltammetrically. A voltammetric measurement can be carried out at a fixed constant potential, i.e. amperometrically, or at a variable potential. During subtraction, a sensitivity factor, set in Example 1, can preferably be determined, which weights the contamination and, if applicable, temperature-corrected interference concentration during subtraction. The nominal temperature is typically 25°C. The analyte charge is determined enzymatically, while the interference charge is determined non-enzymatically. The simple voltammetric measurement is non-specific. The enzymatic-electrochemical measurement provides the selective oranalyte-specific measurement signal, which, however, has a non-specific component due to the voltammetric detection principle and can be corrected by simple voltammetric measurement.

[0051] For the optional temperature correction, the calibration curves can include corresponding pre-stored, i.e., previously recorded, calibration curves between the analyte concentration and the temperature, between the interference concentration and the temperature, and between contamination / ionic conductivity and the temperature. For the contamination correction, i.e., correction for contaminating substances, a contamination, such as hematocrit, is used using previously determined calibration curves / sets of correlation curves between the analyte concentration and the contamination (e.g., hematocrit) and between the interference concentration and the contamination. The pre-stored calibration curves can be stored in a memory / data storage device, which is accessed by a processor implementing the method.

[0052] The term "contamination" and its derivatives, such as "contamination-corrected" or "contamination value," refer to substances that may affect conductivity and influence the determination of the analyte concentration, even though they are not directly electrochemically active themselves. One example is the hematocrit. Another example is contamination by blood components such as leukocytes and their components.

[0053] The first measuring chamber and the second measuring chamber can each preferably comprise a volume of 0.15 pL to 0.3 pL. The temporal sequence can be as follows, for example: (i) The optional temperature measurement can be carried out immediately after filling the measuring chambers and after temporally defined current threshold values ​​of the measuring channels have been exceeded over 50 to 1000 ms at the resistor, (ii) An ionic conductivity measurement for contamination determination, for example hematocrit determination, can follow with a temporal interval of 0.5 s to 1 s, (iii) The measurement for determining the interference concentration in the sample can be carried out between the second and ninth second, (iv) An enzymatic-electrochemical measurement of the analyte concentration can be carried out between the third and tenth second. Thus, the entire measuring process is completed, for example, in approximately 10 s, or at least less than 60 s.The determination of the ionic conductivity of the sample by the four-electrode conductivity arrangement, the determination of the interference charge of electrochemically active substances of the sample by the first three-electrode arrangement and the enzymatic-electrochemical determination of the analyte concentration of the sample by the second three-electrode arrangement is preferably carried out stepwise within a measuring interval of 2 s to 60 s, preferably between 8 and 12 s.

[0054] The described method is particularly suitable for disposable sensors intended to provide clinically relevant and reliable concentration values. The system is particularly suitable for the determination of D-lactate in a patient sample, such as synovial fluid, cerebrospinal fluid, urine, or blood.

[0055] A further technical advantage in some embodiments is that, by using a temperature measuring resistor applied to the carrier, i.e., integrated, the current temperature can be measured with high precision and without susceptibility to interference compared to the prior art (see the above explanations). This precisely measured ambient temperature can then also be used to correct the temperature-dependent enzymatic-electrochemical analyte measurement, as well as to eliminate temperature-dependent interference caused by contamination, such as hematocrit, and electrochemically active substances. The influence of contamination, such as hematocrit or other conductivity-altering cellular components, and interfering electrochemically active substances, is thus corrected more systematically and adequately than with known technical solutions.This allows contamination-corrected concentration values ​​for non-contaminated samples, such as synovial fluid or CSF, to be calculated from contaminated sample material. Because the analyte concentration is determined in a contamination-corrected manner, the disposable sensor is particularly suitable for emergency situations or for quickly required analyte determinations. In contrast to clinical settings, where, for example, contamination can be removed from the sample matrix prior to automated analyte measurements by filtration or centrifugation, which require both a significantly larger sample volume and greater time and equipment expenditure to prepare the sample material, a very small sample volume is used in these application areas, and the analyte value can be measured in seconds in a clinically relevant and reliable manner, even under field conditions.

[0056] To achieve a more accurate determination of the analyte concentration, the method may comprise correcting the analyte concentration and the interference concentration to a contamination-corrected analyte concentration and a contamination-corrected interference concentration using pre-stored calibration curves and the previously determined contamination value, respectively.

[0057] This allows the ionic conductivity measurement to be corrected for ionically conductive components in the sample, particularly for a dissociated analyte or dissociated interference concentrations, such as endogenous or exogenous metabolites. The contamination level, for example, the hematocrit level, is thus determined more accurately, allowing the analyte concentration to be determined, which is dependent on this determined contamination level, and the interference concentration to be subtracted. Calibration curves are to be determined between the ionic conductivity / contamination level and the analyte concentration, or between the ionic conductivity / contamination level and the interference concentration.The analyte-corrected conductivity measurement can be determined using a previously determined non-linear calibration curve between hematocrit and ionic conductivity of the sample to determine the contamination level.

[0058] The optional temperature measuring resistor is preferably a meander conductor structure applied to the sensor carrier. The meander conductor structure allows a sufficient resistance length to be generated on a small carrier surface. This allows measurable temperature-related resistance changes to be recorded, requiring only a small portion of the sensor carrier. Furthermore, due to its small surface area, the meander conductor structure can be positioned close to the measuring chambers, allowing for precise and interference-free temperature determination. This, in turn, improves the accuracy of temperature compensation, allowing the analyte concentration to be determined with greater precision. The meander conductor structure preferably has a resistance between 100 Ω and 2000 Ω and a temperature coefficient between 0.4 Ω / °C and 0.7 Ω / °C for temperature-dependent resistance measurement.

[0059] The meandering conductor structure is preferably positioned adjacent to the measuring chambers. This allows for particularly good temperature measurement at the measuring chambers. This, in turn, improves the accuracy of temperature compensation, allowing the plasma-related analyte concentration to be determined with greater accuracy.

[0060] The meander conductor structure is preferably positioned in a carrier section which, relative to the sample receiving area, has a length that is less than one-third, preferably less than one-quarter, and even more preferably less than one-fifth of the total length of the carrier. This ensures that the temperature is measured close to the measuring chambers. Furthermore, by connecting a measuring device, heat sources can have a reduced influence on the temperature measurement. This allows the temperature in the measuring chambers to be measured more accurately. This, in turn, improves the accuracy of the temperature compensation, so that the analyte concentration to be determined can also be determined with greater accuracy.

[0061] Preferably, determining the ambient temperature comprises determining a first temperature after or during filling the measuring chambers; determining a second temperature after determining the contamination level, the analyte charge, and / or the interference charge of electrochemically active substances in the sample; and determining the ambient temperature by arithmetic averaging of the measured temperatures. This allows incorrect temperature determination due to temperature drift during the measurement process to be minimized or reduced by averaging. This, in turn, improves the accuracy of temperature compensation, so that the analyte concentration to be determined can also be determined with greater accuracy, since the undesirable temperature drift is compensated or averaged out. The two temperature-dependent resistance measurements on the meandering conductor structure can be performed over a period of 50 to 1000 ms.

[0062] The first measuring chamber preferably comprises the first three-electrode arrangement and the four-electrode conductivity arrangement; the second measuring chamber preferably comprises the second three-electrode arrangement. The first three-electrode arrangement is preferably a voltammetric electrode arrangement. The second three-electrode arrangement is preferably a voltammetric electrode arrangement.

[0063] The method further comprises switching the electrodes in the first measuring chamber between the first three-electrode arrangement for voltammetric measurement and the four-electrode conductivity arrangement for measuring ionic conductivity by means of an integrated or reversibly connected analog switch array. This makes it possible for one measuring chamber to be designed for two measured variables, namely ionic conductivity and interference concentration. Existing electrodes can thus be used twice. This not only saves electrode material but also enables a more compact measuring space, which can be more accurately measured temperature-wise by the meander conductor structure. This, in turn, improves the accuracy of temperature compensation, so that the analyte concentration to be determined can also be determined with greater accuracy.

[0064] Preferably, the first three-electrode and four-electrode conductivity arrangement comprises a reagent coating comprising a redox mediator, and wherein the second three-electrode arrangement has a reagent coating comprising an oxidoreductase or other catalytically active proteins and a redox mediator.

[0065] In some embodiments, the second three-electrode assembly comprises a reagent coating comprising an oxidoreductase and a redox mediator.

[0066] The total protein content is preferably determined by the amount of an oxidoreductase or by an oxidoreductase and one or more additional catalytically active proteins. The oxidoreductase preferably comprises oxidases, peroxidases, and / or cofactor-dependent dehydrogenases. The catalytically active proteins preferably comprise hydrolases, proteases, and esterases. The redox mediator preferably comprises a redox-active metal complex, a quinoid redox dye, or an organometallic compound.

[0067] The oxidoreductase is preferably a D-lactate dehydrogenase, preferably an NAD-independent D-lactate dehydrogenase.

[0068] In one embodiment of the method, the sample is incubated with a reagent mixture comprising a buffer solution, a redox mediator, and a D-lactate dehydrogenase, preferably an NAD-independent D-lactate dehydrogenase. Optionally, the reagent mixture may comprise a stabilizing additive. Examples of such additives that can serve to stabilize proteins are glycerol, mannitol, sorbitol, histidine, arginine, poloxamers, sucrose, glycine, trehalose, and mixtures thereof.

[0069] In one embodiment, the method further comprises measuring a negative control with a reagent mixture comprising a buffer solution, a redox mediator, and optionally an additive for stabilization.

[0070] For the temperature-dependent resistance measurement of the meander conductor arrangement, a two-electrode conductivity arrangement is preferably used, and even more preferably a four-electrode conductivity arrangement. When measuring the ionic conductivity, a current injection of between 100 pA and 750 pA is preferably used.

[0071] For the four-electrode conductivity arrangement, an alternating voltage without a direct voltage component is preferably applied, wherein the alternating voltage is rectangular, triangular or sinusoidal with a preferred frequency between 100 Hz and 5000 Hz. Preferably, each electrode can be switched off in a defined manner via an analog switch array or connected together with other electrodes.

[0072] Preferably, the method comprises measuring the ionic conductivity of the sample by the four-electrode conductivity arrangement, voltammetrically determining the interference charge of electrochemically active substances of the sample by the first three-electrode arrangement, and enzymatically electrochemically determining the analyte charge of the sample by the second three-electrode arrangement step by step within a measuring interval of 8 s to 20 s, preferably between 8 s and 11 s. Furthermore, the short-term measuring process can reduce temperature interference effects.The time distribution after filling the measuring chambers with sample can be as follows: (i) The measurement of the meander resistance to determine the temperature can be carried out over 50 ms to 1000 ms, (ii) An ionic conductivity measurement to determine contamination can follow with a time interval of 0.5 s to 1 s, (iii) The voltammetric measurement to determine the interference concentration in the sample in the first measuring chamber can be carried out between the second and ninth second, (iv) An enzymatic-electrochemical measurement of the analyte concentration can be carried out between the third and tenth second, (v) A second measurement of the meander resistance to determine the temperature can be carried out over 50 ms to 1000 ms.

[0073] In a further aspect of the invention, a sensor for determining an analyte concentration in a sample is described. This sensor comprises at least two measuring chambers, which can be filled with a sample via a sample receiving area of ​​the sensor, wherein at least two measuring chambers are formed on a carrier and comprise a first three-electrode arrangement, a four-electrode conductivity arrangement, and a second three-electrode arrangement. The sensor can optionally have a temperature measuring resistor applied to the carrier for determining an ambient temperature. The sensor further comprises at least one processor unit, which is reversibly connected to the sensor via an electrical contact or is integrated on the carrier, and is configured to do the following:

[0074] Controlling the four-electrode conductivity array to determine the ionic conductivity of the sample;

[0075] - controlling the first three-electrode arrangement to determine an interference charge of electrochemically active substances of the sample;

[0076] - Controlling the second three-electrode arrangement to determine an analyte charge of the sample enzymatically-electrochemically;

[0077] - Determining an analyte concentration using pre-stored calibration curves and the determined analyte charge;

[0078] - Determining an interference concentration using pre-stored calibration curves and the determined interference charge; and determining a contamination value using pre-stored calibration curves and the determined ionic conductivity;

[0079] - Correcting the analyte concentration and the interference concentration to a contamination-corrected analyte concentration and a contamination-corrected interference concentration using pre-stored calibration curves and the previously determined contamination value;

[0080] Determining the interference-corrected analyte concentration by subtracting the interference concentration from the analyte concentration. The processor unit is preferably a processor or a microcontroller. The pre-stored calibration curves can be stored in a memory / data storage device to which the processor unit is operatively connected.

[0081] This results in the same advantages as described for the above procedure, to which reference is hereby made.

[0082] The analyte concentration is preferably the concentration of D-lactate in the sample.

[0083] In one embodiment, the sensor further comprises an applied temperature measuring resistor for determining an ambient temperature. The sensor is preferably configured to control the temperature measuring resistor in order to determine the ambient temperature after the measuring chambers have been filled. The temperature measuring resistor is preferably a meander conductor structure applied to the sensor carrier. The meander conductor structure is preferably positioned adjacent to the measuring chambers.

[0084] The meander ladder structure is preferably positioned in a carrier section which, relative to the sample receiving area, has a length which is less than one third, preferably less than one quarter, even more preferably less than one fifth of the total length of the carrier.

[0085] Preferably, the first measuring chamber comprises the first three-electrode arrangement and the four-electrode conductivity arrangement, and wherein switching of electrodes in the first measuring chamber between the first three-electrode arrangement for voltammetric measurement and the four-electrode conductivity arrangement for measuring the ionic conductivity can be carried out by means of an integrated or reversibly connected analog switch array.

[0086] Further preferred embodiments of the sensor can also be derived from the above method in a corresponding manner.

[0087] Short description of the characters

[0088] The invention is explained in more detail below using an exemplary embodiment and the accompanying drawings. The figures show:

[0089] Figure 1 shows a sensor according to a preferred embodiment of the invention; Figure 2 shows a schematic representation of a sensor according to a preferred embodiment of the invention;

[0090] Figure 3 is a schematic representation of a flow chart according to the invention;

[0091] Figure 4 shows an exemplary pre-stored calibration curve of the meander resistance measurement value as a function of the ambient temperature;

[0092] Figure 5 shows exemplary pre-stored calibration curves for resistance measurement (1 / ionic conductivity) as a function of the hematocrit of the sample at ambient temperatures of 15°C, 25°C and 45°C in the first measuring chamber;

[0093] Figure 6 shows exemplary pre-stored calibration curves for charge values ​​recorded as a function of the lactate concentration in the second measuring chamber at temperatures between 5°C and 45°C;

[0094] Figure 7 shows an exemplary pre-stored set of curves generated from calibration curves of Figure 6 by plotting the lactate concentration against the temperature for different charge values ​​measured in the second measuring chamber;

[0095] Figure 8 shows exemplary pre-stored calibration curves for charge values ​​as a function of lactate concentration at hematocrit concentrations of 0% (plasma), 19%, 45% and 70% measured in the second measuring chamber;

[0096] Figure 9 shows an exemplary pre-stored set of curves generated from calibration curves in Fig. 8 by plotting the lactate concentration against the hematocrit value at different charge values ​​measured in the second measuring chamber;

[0097] Figure 10 exemplary pre-stored calibration curves for charge measurements against a dilution series of a model solution of 0.9 mM ascorbic acid, 1.03 mM paracetamol and 1.4 mM uric acid, measured in the first and second measuring chamber;

[0098] Figure 11 shows a calibration curve of a D-lactate sensor;

[0099] Figure 12 Measurements of the D-lactate sensor signal at different amounts of enzyme (lactate dehydrogenase);

[0100] Figure 14 shows the D-lactate sensor behavior in the presence of different buffer solutions;

[0101] Figure 15 shows the D-lactate sensor signal for different D-lactate concentrations as a function of incubation time; Figure 16 shows a bar chart of D-lactate sensor signals as a function of incubation time;

[0102] Figure 17 shows the signal-to-noise behavior of the D-lactate sensor depending on different temporal measurement segments;

[0103] Figure 18 shows the signal-to-noise behavior of the D-lactate sensor at different polarization voltages;

[0104] Figure 19 the D-lactate sensor signal after switching on the polarization voltage in the 1st second;

[0105] Figure 20 dose-response curves for a D-lactate dehydrogenase sensor assay according to the invention;

[0106] Figure 21 shows the determination of the detection limit of a D-lactate sensor according to the invention in buffer and artificial synovial fluid;

[0107] Figure 22 shows the determination of the detection limit for D-lactate in single-strength artificial synovial fluid;

[0108] Figure 23 shows the sensor signal of the D-lactate sensor at the indicated concentrations of various substances;

[0109] Figure 24 shows a dose-response curve of the D-lactate sensor with real patient samples; and

[0110] Figure 25 shows the determination of the detection limit for D-lactate using the measurements in Figure 24.

[0111] Detailed description of the invention

[0112] Fig. 1 shows a sensor 100 according to a preferred embodiment of the invention. The sensor 100 is, in particular, a single-use sensor, also referred to as a disposable sensor.

[0113] The sensor 100 comprises a planar carrier 1 extending along a longitudinal axis. Appropriate sensor components are applied to the carrier 1. The carrier material is preferably a plastic, for example, PET (polyester). The carrier 1 can have a thickness of 0.25 mm, for example, although the invention is not limited thereto.

[0114] The sensor 100 further comprises a sample receiving area 17 at a sample receiving end of the carrier 1, via which a whole blood sample, for example a capillary blood sample, can be guided into measuring chambers 2, 3. The measuring chambers 2, 3 have a common sample receiving area 17 and run parallel to one another. Furthermore, the measuring chambers 2, 3 are applied to the carrier 1 in a liquid-tight manner. An electrically insulating lacquer layer 14 can be printed underneath, which contains two parallel recessed measuring windows and thus defines or delimits the associated electrode arrangements with respect to their areas. Furthermore, the meander structure of the temperature measuring resistor (11) can be recessed. The measuring chambers 2, 3 can preferably be designed for a filling volume between 150 nL and 300 nL. Furthermore, venting channels 18a, b can be provided.

[0115] In the enlarged section of Figure 1, electrodes 4, 5, 6, 7a, 7b, 8, 9, 10 are also visible on the carrier 1. In addition, a temperature measuring resistor 11 is provided on the carrier

[0116] I of the sensor 100 to measure the ambient temperature. Positioning it on the carrier 1 makes the temperature measurement less susceptible to interference, and the distance to the measuring chambers can be kept small. The temperature measuring resistor 11 is designed as a meandering conductor structure / meander-shaped conductor structure. This allows for the detection of measurable temperature-related resistance changes with sufficient accuracy, requiring only a small surface area of ​​the sensor's carrier 1.

[0117] The meander conductor structure 11 is positioned (immediately) adjacent to the measuring chambers 2, 3 and thus in close proximity to the measuring chambers 2, 3. This allows the temperature of the measuring chambers 2, 3 to be measured particularly well and reliably. This, in turn, improves the accuracy of the temperature compensation of the required measured variables, which are described in more detail in Figure 3.

[0118] The meander conductor structure 11 is positioned in a support section D of the support 1, which has a length that is less than one-third, preferably less than one-fifth, of the total length L of the support 1. As a result, heat interference sources, such as the connection of a measuring device and its operation, can have a reduced interference influence on the temperature measurement, so that the temperature in the measuring chambers 2, 3 can be measured more accurately.

[0119] The meander conductor structure 11 can be completely covered by the insulating varnish 14 or preferably also left uncoated. Furthermore, supply lines 12 to the electrodes 4, 5, 6, 7a, 7b, 8, 9, 10 and to the meander conductor structure 11 are provided, which enable the operation of the electrodes 4, 5, 6, 7a, 7b, 8, 9, 10 and the meander conductor structure

[0120] II. Furthermore, electrical contacts 13, in particular contact surfaces, are applied to the carrier 1 at the end opposite the sample receiving area 17. A measuring device, or in particular a processor 50 of a measuring device, can be reversibly connected to the sensor 100 via the electrical contacts 13, as schematically illustrated in Figure 2.

[0121] In a special embodiment, the structures can be manufactured as follows. After a sputtering process, in which a thin layer of an inert metal, for example a 50 nm thick gold layer, is applied, a first three-electrode arrangement 31 with working electrode AE1 4, counter electrode GE1 5, and reference electrode RE1 6, including two additional voltage-tapping measuring electrodes 7a, b (four-electrode arrangement 33), is ablated using a laser. In parallel, a second three-electrode arrangement 32 with working electrode AE2 8, counter electrode GE2 9, and reference electrode RE2 10 is ablated. In the immediate vicinity, the meander conductor structure 11, including supply lines 12 and electrical contacts 13, can be structured by ablation.

[0122] The electrical contacts 13 can form contact surfaces that serve to establish secure electrical contact with a measuring device. The insulating lacquer layer 14 can contain two parallel recessed windows, each of which defines an arrangement of electrode surfaces. The windows can correspond to the plane-parallel microfluidic measuring chambers 2, 3. Furthermore, a third recess can be present in the insulating lacquer layer 14, which defines the meander structure 11.

[0123] The measuring chambers 2, 3 comprise a first three-electrode arrangement 31, a four-electrode conductivity arrangement 33 and a second three-electrode arrangement 32. These are described in more detail below using a preferred embodiment.

[0124] In this embodiment, the first measuring chamber 2 contains the electrodes 4, 5, 6, 7a, b. The electrodes 4, 5, 6, 7a, b are each connected to an analog switch array 55. With the aid of this analog switch array 55, controlled by a processor 50, the electrodes 4, 5, 6, 7a, b can be used for dual functions, so that they can be connected either to a first voltammetric three-electrode arrangement 31 with a working electrode AE1 4, a counter electrode GE1 5, and a reference electrode RE1 6, or to a four-electrode conductivity arrangement 33 with two current-supplying electrodes 5, 6 and two voltage-tapping electrodes 7a, b. This makes it possible to realize a dual use of the first measuring chamber 2 and thus to realize a compact measuring area with resistance temperature meter 11 on the carrier 1 and in the vicinity of the measuring chambers 2,3.Electrodes 4, 5, 6, 7a, and b are coated with a reagent layer consisting of redox mediator, electrolyte-forming ions, and detergents. The redox mediator content can amount to between 10 and 20 pg of the reagent load.

[0125] The first three-electrode arrangement 4, 5, 6 in the first measuring chamber 2 determines the interference concentration of electrochemically active substances in the sample. The four-electrode conductivity measuring arrangement 5, 6, 7a, b is used in the first measuring chamber 2 to determine a contamination level of the sample based on the ionic conductivity.

[0126] The second three-electrode arrangement 32 in the second measuring chamber 3 further comprises a second three-electrode arrangement 32 with a working electrode AE2 8, counter electrode GE2 9, and reference electrode RE2 10, and determines the analyte concentration in the sample using enzymatic voltammetry. The electrodes can be coated with a reagent system consisting of an oxidoreductase and, optionally, one or more additional catalytically active proteins, electrolyte-forming ions, and a redox mediator. The total protein quantity is preferably 50 pg to 80 pg, and the electron mediator portion makes up 40 pg to 80 pg of the reagent load. Oxidases, peroxidases, or cofactor-dependent dehydrogenases are preferably used as oxidoreductases. Additional catalytically active proteins that may be used are hydrolases, proteases, and esterases. The preferred redox mediator is a redox-active metal complex, a quinoid redox dye or an organometallic compound.

[0127] For the meander conductor structure 11, also called a meander-shaped conductor track, applied in the immediate vicinity of the measuring chambers 2, 3, i.e., directly adjacent to the measuring chambers 2, 3, two supply lines for current input and two supply lines for voltage tapping are arranged, thus enabling a four-wire resistance measurement. The meander conductor structure 11 and supply lines 13 can be covered with the insulating lacquer layer 14. The meander structure is preferably left uncoated by the lacquer layer. The meander resistance can be between 100 Ω and 2000 Ω at 25°C.

[0128] The lead ends of both the conductivity and resistance measuring electrodes, as well as the lead ends of the two voltammetric three-electrode arrangements, are formed as electrical contacts 13 at the end of the carrier 1 of the sensor 100 opposite the sample receiving area 17. As schematically illustrated in Figure 2, these allow contact with a measuring device or with a processor 50. The processor 50, or the measuring device comprising the processor 50, carries out the method according to the invention, which is described in more detail in Figure 3.

[0129] The pre-stored calibration curves can be stored in a memory / data storage 60 to which the at least one processor 50 is operatively connected.

[0130] For example, the measuring device provides the required operating voltages for the respective measuring channels, controls the measuring procedure, processes the measuring signals, displays the measurement results, and stores them, for example, in memory 60. Such a measuring device can have voltammetric measuring channels, a resistance measuring channel, and a measuring channel for measuring ionic conductivity. Furthermore, the measuring device can have an analog switch array 55 for functionally assigning the electrodes. The polarization voltage for the enzymatic amperometric lactate measurement is, for example, +175 mV, and the polarization voltage for the amperometric measurement of the electrochemically active substances is +400 mV, respectively, against the internal reference electrode RE2 or RE1 of the sensor 100.The current input for the four-electrode resistance measurement of the meander conductor structure 11 can be operated with 100 pA and the ionic conductivity measurement for hematocrit determination can be carried out with a rectangular, triangular or sinusoidal alternating voltage (f = 1000 Hz) with an amplitude of 500 mV.

[0131] The method according to the invention is described below by way of example with reference to Figure 3, which can be carried out by a processor 50 integrated on the carrier 1 or electrically contacted.

[0132] Fig. 3 shows a schematic representation of a flow chart according to the invention. The determination of a contamination-corrected analyte value, for example, for lactate, glucose, etc., which is corrected for interfering influences such as hematocrit and electrochemically active substances in the sample as well as for the influence of the ambient temperature, is carried out using the sensor 100 according to the invention and the method as described in more detail in the following embodiment.

[0133] In a first step, not explicitly shown here, the at least two measuring chambers 2, 3 (see Fig. 1) are filled with a sample (synovial fluid, cerebrospinal fluid, urine, blood, capillary blood) via the sample receiving area 17 of the sensor. The two measuring chambers 2, 3 are formed on the carrier 1 of the sensor and, as already described above, comprise a first three-electrode arrangement 31, a four-electrode conductivity arrangement 33, and a second three-electrode arrangement 32.

[0134] In one step, an ambient temperature S100 is determined after the measuring chambers 2, 3 have been filled by a temperature measuring resistor 11 applied to the carrier. For example, the temperature can be determined immediately after the measuring chambers 2, 3 have been filled and time-defined current thresholds of the voltammetric measuring channels have been exceeded. The four-wire resistance measuring channel can perform a first temperature-dependent resistance measurement over 50 to 1000 ms on the meander conductor structure 11. For current generation by the meander conductor structure 11, 100 to 500 pA of direct current can be provided by the measuring channel. The measured voltage drop across the meander conductor structure 11 at a predetermined current serves to determine the resistance and can be temporarily stored, for example in memory 60 (see Fig. 2). An ambient temperature can then be determined from the resistance.

[0135] The ambient temperature can be determined from the linear relationship between temperature (T) and ohmic resistance (R) of the resistance measurement on the meandering conductor 11 according to equation (1) as the mean value between the resistance values ​​measured at the beginning and end of the measurement procedure. This can mitigate temperature deviations caused by temperature drift during the measurement process.

[0136] T — kyemp RlVIä + RlVIäO (1 )

[0137] Here, kTem - the temperature coefficient (temperature sensitivity) of the meander conductor

[0138] R Mä - the measured meander resistance

[0139] RMäo - the meander resistance at 0°C

[0140] The temperature coefficient (kTemp) was previously determined experimentally using a calibration curve recorded between 0°C and 50°C as a function of the resistance value of the meander conductor T = f (R Ma) and can be between 0.3 and 0.8 °C / Q. The calibration curve, as well as all other pre-stored calibration curves, can be stored in memory 60.

[0141] In a further step, the ionic conductivity S110 of the sample is determined using the four-electrode conductivity arrangement 33.

[0142] Furthermore, the temperature-corrected ionic conductivity S210 is determined using pre-stored calibration curves, the measured ambient temperature, and the determined ionic conductivity. Pre-stored calibration curves and the ambient temperature determined by the temperature measuring resistor 11 applied to the carrier 1 are used.

[0143] For example, in the first measuring chamber 2, the counter electrode GE1 5 and reference electrode RE1 6 for current supply and two voltage tapping electrodes 7 a, b for tapping the voltage drop can be connected to the four-electrode conductivity measuring channel via an analog switch array 55.

[0144] The four-electrode conductivity arrangement 33 can preferably be supplied with a sinusoidal alternating voltage with a frequency of 100 Hz to 1000 Hz and an amplitude between 100 mV and 1000 mV. The resulting voltage drop, which depends on the ionic conductivity of the first measuring chamber, is corrected for the phase angle if necessary, and the resistance of the meander conductor structure 11 is determined. The resistance value can be measured and temporarily stored, for example, within 0.25 s to 1.0 s after the end of the temperature-dependent resistance measurement.

[0145] A (preliminary) determination of the hematocrit concentration (Hct) is performed according to step S210 as follows: The resistance or ionic conductivity (G) determined in the first measuring chamber depends essentially on the hematocrit, the temperature, and, if applicable, the concentration of a dissociated analyte or another dissociated endogenous or exogenous metabolite of the sample. The relationship between the measured resistance (R = 1 / G) and hematocrit (Hct) is described by a parabolic equation (2):

[0146] Rrict = fr (Hct) = ki * concHet 2 + k2* conc Hc t + k3T e {1°C, 5°... 45°C}

[0147] (2)

[0148] Here, the following means:

[0149] Rnct- hematocrit-dependent resistance measurement value, determined from the ionic conductivity measurement in the first measuring chamber concHet - hematocrit concentration k3- resistance of the plasma value (Hct-free sample) ki and k2- cell constants f T (Hct) - temperature-dependent parabolic family

[0150] A required semi-parabolic family f T (Hct) is determined using

[0151] For blood samples with hematocrit values ​​between, for example, 0% and 70%, or for samples (e.g., synovial fluid, cerebrospinal fluid, urine) with varying degrees of blood contamination, the resistance values ​​are determined at temperatures between, preferably, 1°C and 45°C in at least seven stages, as an example. The previously determined ambient temperature is then used to assign the hematocrit-dependent resistance measurement value or hematocrit value in the parabolic family.

[0152] The calculation of the temperature-corrected ionic conductivity can be performed by rearranging equation (2) as follows: concHet

[0153] For resistance values ​​that lie between two isothermal Hct semi-parabolas of the array, a linear interpolation of the Hct value is performed based on the measured temperature. This results in a temperature-corrected hematocrit value.

[0154] In a further step of the method, the voltammetric determination of an interference concentration S120 of electrochemically active substances in the sample is carried out by the first three-electrode arrangement 31. The analog switch array 55 (see Fig. 2) can separate the four electrodes 5, 6, 7a, 7b, 8 from the conductivity measuring channel, and the counter electrode GE1 5 and reference electrode RE1 6 can be connected together with the working electrode 8 as a potentiostatic first three-electrode arrangement 31 in the first measuring chamber 2. The electrodes can be connected according to the first voltammetric measuring channel for the amperometric measurement of the electrochemically active substances, with the working electrode being subjected to a polarization voltage between 100 mV and 500 mV.

[0155] In a further step, the enzymatic-electrochemical determination of an analyte concentration or analyte charge S130 of the sample is carried out by the second three-electrode arrangement 32. In parallel or in series, the second potentiostatic three-electrode arrangement consisting of the working electrode AE2 8, the counter electrode GE2 9, and the reference electrode RE2 10, which is arranged in the second measuring chamber 3, is activated by connecting the second voltammetric measuring channel for the enzymatic-electrochemical determination of the analyte concentration. The provided polarization voltage is preferably between 100 mV and 500 mV.

[0156] The evaluation can preferably be carried out in the following way: The measured signal currents for measuring the analyte concentration and electrochemically active substances are essentially described by the Cotrell equation (4): Herein mean:

[0157] I - current z - number of transferred electrons F - Faraday constant (96485 As / mol) D - diffusion constant

[0158] A - Electrode surface of the working electrode t - Time c - Initial concentration of the reacted substance

[0159] The measuring currents are preferably integrated over time-defined intervals and the resulting charge values ​​are proportional to the analyte concentration or the interference concentration (the concentration of interfering substances) and, in addition, depend on the temperature and any cellular components (e.g., hematocrit).

[0160] Furthermore, a temperature-corrected analyte concentration S230 and a temperature-corrected interference concentration S220 are determined using pre-stored calibration curves, the ambient temperature determined by the temperature measuring resistor applied to the carrier, and the analyte charge and interference charge determined. This is described in more detail in the following example.

[0161] For this purpose, previously recorded sets of curves are used according to the following function sets

[0162] QTA = fr (concAnalyte) = k Ai concAnalyt 2 + k A 2 conc Analyte + k A 3, T e {1°C, 5°... 45°C} (5) and

[0163] Q T | = fr (concmterf.) = kn concAnalyt 2+ k|2 conc analyte + k|3, T e {1 °C, 5°... 45°C} (6) and which are recorded in at least seven temperature steps, depending on the analyte or interferent concentration of the samples, for example for temperatures between 1°C and 45°C.

[0164] Here, the following means:

[0165] QTA - temperature-dependent analyte charge values ​​of the curve sets

[0166] QTI - temperature-dependent interference charge values ​​of the curve families T - ambient temperature (curve family parameters) cone Anaiyt - analyte concentration cone interf. - interfering substance concentration kAi , kA2 - analyte sensitivity kA3 - charge value at c Anaiyt = 0 mM kn, ki2 - interfering substance sensitivity kn - charge value at Cintf = 0 mM

[0167] The function families (5) and (6) can then be rearranged according to the concentration and the concentration values ​​can be plotted against the temperature using a series of closely spaced charge values ​​(n), resulting in a family of n curves whose function families are each described by a general quadratic equation (7), (8):

[0168] COnC analyte Tcorr = f QTA (T) = AT1 T 2 + AT2 T + AT3 QTA £ {0, 5 pC, 10 pC ... 60 pC} (7) cone Interf-Tcorr = foTI (T) = k|T1 T 2 +k|T2 T + k|T3 QTI e {0, 0.3 pC, 0.6 pC ... 6.0 pC} (8)

[0169] Where: conc interf-Tcorr - temperature-corrected analyte value conc interf-Tcorr - temperature-corrected interferent value

[0170] QTA - temperature-dependent analyte charge values ​​(curve family parameters)

[0171] QTI - temperature-dependent interference charge values ​​(curve family parameters)

[0172] T - Ambient temperature kATi, kAT2 - Temperature sensitivity of the analyte measurement kAT3 - Analyte concentration value at T = 0°C km, kiT2 - Temperature sensitivity of the interference measurement kiT3 - Interference concentration value at T = 0°C

[0173] The intersection point of the measured charge value and temperature, which lies either on one of the calibration curves of the family of curves or between two adjacent curves, is used to determine the curve that is closest to this intersection point.

[0174] If the measured charge value for a concentration does not lie on a concentration isoline but between two concentration lines, a linear interpolation is performed.

[0175] The functional equation of this curve is used to determine the corresponding concentration for the analyte and the interfering substances for the standard temperature of preferably 25°C.

[0176] In a preferred embodiment, the correction of the temperature-corrected ionic conductivity value to an analyte- and temperature-corrected conductivity value (S240) can be performed using pre-stored calibration curves and the determined temperature-corrected analyte concentration, and / or using pre-stored calibration curves and the determined temperature-corrected interference concentration.

[0177] This is described in more detail below using an example of a blood sample and the biomarker lactate. The temperature-corrected raw values ​​of the analyte and interference concentrations are used to correct the hematocrit value according to Gig. (9) to compensate for the change in conductivity or resistance due to the presence of a dissociated analyte or an electrochemically active interfering substance in the sample, which can lead to an erroneous hematocrit determination.

[0178] The basis for this is a regression line according to equation (7), which was previously plotted as a function of the conductivity or resistance of a dissociated substance concentration in the form of the analyte or an endogenous or exogenous metabolite:

[0179] Temperature 25°C ( — 1 / GHC — f (Cdis) - 3dis Cdis + bdis (9)

[0180] Here, the following means:

[0181] Rrict 25°C - hematocrit-dependent resistance value, determined from the ionic conductivity Gnct in the first measuring cell at 25°C

[0182] Cdis - Concentration of a dissociated substance concentration

[0183] -adis - negative increase of the resistance Rnct bdis - resistance Rnct in the absence of a dissociated substance concentration

[0184] The known substance concentration that is to be expected in the blood sample and causes the largest negative increase is used to correct the resistance value to determine the Hct according to equation (10): RlHct-corr — adis * COFICRW + Rnct 25°C (10)

[0185] Where: concRw - raw value of the concentration of a dissociated substance concentration adis - increase in resistance Rnct 25°c at 25°C as a function of the concentration of the dissociated substance concentration

[0186] Rnct-corr - hematocrit-dependent resistance measurement value, corrected for the reduction of the resistance value due to the presence of a dissociated substance concentration at 25°C.

[0187] In a further step, the temperature-corrected analyte concentration and the temperature-corrected interference concentration are corrected to a plasma-related hematocrit and temperature-corrected analyte concentration (S330) and a plasma-related hematocrit and temperature-corrected interference concentration (S320) using pre-stored calibration curves and the previously determined temperature-corrected hematocrit value or the analyte and temperature-corrected hematocrit value.

[0188] In the following, the determined hematocrit value, which in turn is corrected for temperature and optionally for the presence of ionically conductive substances, is used to correct the analyte concentration value and the concentration value of electrochemically active substances (interference substance concentration). In other embodiments, only the temperature-corrected hematocrit value is used, for example, if no significant additional contribution from ionically conductive substances is expected. For this purpose, a previously recorded set of straight lines according to equations (11) and (12) is used. These are plotted against a reference system as a function of the hematocrit, for example, for a range from 0% to 70% or the corresponding Hct-dependent resistance value in at least 7 steps at a temperature of 25 °C:

[0189] Q Hct Analyt — fhlct (COnC Analyte) — 3A COnC Analyte Temp corr + b AHct e {0%, 20%...70% (11) and

[0190] Q Hct Interf — fhct (COnC Interf) — 3lnt COnC Interf Temp corr + bint Hct e {0%, 20%...70%} (12) Where:

[0191] Het - Hematocrit value (line parameter of the temperature-corrected analyte concentration curves and interferent concentration curves)

[0192] ÜHct-Anaiyt - hematocrit-dependent charge values ​​of the temperature-corrected analyte concentration values

[0193] Q Hct-interf - hematocrit-dependent charge values ​​of the temperature-corrected analyte concentration values ​​of the interferent concentration values

[0194] COnC Analyt Temp corr temperature-corrected analyte concentration values

[0195] COnC Interf Temp corr temperature-corrected interference concentration values ​​a A - Slope (analyte sensitivity) b A - Constant, charge value at cAnaiyt = 0 mM a int - Slope (interfering substance sensitivity) b int - Constant, charge value at cinterfer. = 0 mM

[0196] Based on the functional equations (11), (12), the respective concentration values ​​can be determined from the measured charge values ​​QHct-Anaiyt and QHct-interf and plotted against the hematocrit concentration using a series of closely spaced charge values ​​(n), resulting in a family of n curves, each of which is described by a general quadratic equation (13), (14):

[0197] COnC Analyt-T+Hct corr = foTHct-A (Hct) = kyHct AI Hct 2 +kyHct A2 Hct + kyHct A3 QTHct-A £ (OpC, 5 pC, 10 pC....6O pC} (13)

[0198] COnC|nterf-T+Hct corr — foTHct-l (Hct) - kTHct U Hct 2+k THct I2 Hct + k THct I3 QTHct-l £ (OpC,

[0199] 0.3 pC, 0.6 pC....6.0 pC} (14)

[0200] Where: conc Anaiyt-T+Hct corr - temperature and hematocrit corrected analyte value conc interf-T+Hct corr - temperature and hematocrit corrected interferent value

[0201] Q THct-A - Temperature-compensated hematocrit-dependent analyte charge (curve family parameter)

[0202] Q THct-l - Temperature-compensated hematocrit-dependent interference discharge of the (curve family parameters)

[0203] Hct - temperature-corrected hematocrit value k THcti AI , kTHct A2 - temperature-compensated hematocrit sensitivity of the analyte measurement k THcti n , k THct i2 - temperature-compensated hematocrit sensitivity of the interference measurement k THct A3, k THct i3 - concentration values ​​at Hct = 0 % (plasma value)

[0204] The intersection point of the measured charge value and hematocrit, which lies either on one of the curves of the family of curves or between two adjacent curves, is used to determine the curve that is closest to this intersection point.

[0205] If the measured charge value for a concentration does not lie on a charge isoline but between two charge lines, a linear interpolation is performed.

[0206] The functional equation of this curve is used to calculate the corresponding concentration for the analyte and the interferents for a hematocrit concentration of 0% (plasma value).

[0207] In a further step of the method, the analyte concentration is determined (S430) by subtracting the hematocrit and temperature corrected interference concentration from the hematocrit and temperature corrected analyte concentration.

[0208] This is described in more detail below. In this evaluation step, the influence of electrochemically active substances on the determination of the analyte concentration value can be corrected by subtraction and by considering a sensitivity factor according to equation (15):

[0209] COnC Analyt_T+Hct+lnt -corr — COnC Analyt_T+Hct-corr C int_T+Hct corr. *K s (15)

[0210] These are: conc Anai y t_T+Hct+int -corr plasma-related analyte concentration, corrected for the influence of ambient temperature, hematocrit and electrochemically active substances conc Anai y t_T+Hct-corr Plasma-related analyte concentration, corrected for the influence of ambient temperature and hematocrit C int_T+Hct corr. Plasma-related total concentration of electrochemically active substances, corrected for the influence of ambient temperature and hematocrit

[0211] K sSensitivity factor (quotient of sensitivity of the amperometric analyte measuring system and the amperometric interference measuring system against electrochemically active substances)

[0212] In special versions, K s =1 if a similar reagent composition is assumed in both measuring chambers.

[0213] The sensor and method according to the invention thus enable correction of the mutually influencing parameters and thus a more precise approach to correcting the analyte value. This is also based on the very precisely determined ambient temperature directly on the carrier and in the vicinity of the measuring chambers.

[0214] In a corresponding manner, the measurement, correction and evaluation can also be carried out without temperature measurement and temperature-based correction.

[0215] Instead of the hematocrit value, other contaminants can be measured in a corresponding manner and used for the measurement value correction.

[0216] Example 1

[0217] In the following Figures 4 to 10, the method described above and the sensor implementing the method are demonstrated using a concrete example with L-lactate as the analyte to be determined in whole blood. For additional steps, reference is made to the procedure described above for Figure 3. The concrete example further supplements the above information.

[0218] For the determination of lactate, the electrodes 4, 5, 6 of the first measuring chamber 2 within the first measuring window can preferably be coated with a reagent layer of redox mediator (20 pg / mL), sodium chloride (2 mM), Tergitol (0.3% V / V) and CMC (0.5% W / V). This enables both a quantitative summary measurement of electrochemically active substances (interference substances) in the sample using the first voltammetric three-electrode arrangement 31 and the hematocrit-dependent ionic conductivity measurement using the four-electrode conductivity arrangement 33. For lactate, the electrodes 8, 9, and 10 within the second measuring chamber 3 are preferably coated with a reagent solution consisting of lactate oxidase (2 pg / mL), sodium chloride (50 mM), CMC (0.5% w / w), Tergitol (0.3% v / v), and ferricyanide (100 pg / mL) as a redox mediator. The total volume of the reagent application is preferably 200 nL.The meander resistance of the temperature measuring resistor 11 is, purely as an example, 560 ohms at 25°C and has a temperature dependence of 0.6 Ω / °C; see Fig. 4.

[0219] The lead ends of both the conductivity and resistance measuring electrodes, as well as the lead ends of the two voltammetric three-electrode arrangements, are formed as contact surfaces 13 at an end of the carrier 1 opposite the sample receiving area 17. These contact surfaces 13 enable contact with a processor 50, see Fig. 2. The processor 50 can be part of a connected measuring device, as described in connection with Figures 1 and 2.

[0220] Immediately after filling the measuring chambers with a capillary blood sample and exceeding time-defined current thresholds of the voltammetric measuring channels, a first temperature-dependent resistance measurement is performed on the four-wire arrangement of the meander conductor structure 11, preferably over 500 ms. 100 pA of direct current is fed into the meander conductor structure 11 via the measuring channel as an excitation signal. The measured resistance value, in this specific case 549.1 Ω, is temporarily stored. In other embodiments, this resistance value can be used directly. This step corresponds to step S100 in Fig. 3.

[0221] After the resistance measurement, the counter electrode GE 5 and reference electrode RE1 6, intended for current input, as well as the two voltage-tapping electrodes 7a, b of the first measuring chamber 2, are connected to the four-electrode conductivity measurement channel of the measuring device via analog switches of the internal analog switch array 55 (see Fig. 2). The working electrode AE1 4 is switched off. The conductivity measurement channel can, for example, provide a rectangular alternating voltage with a frequency of 1000 Hz and an amplitude of 100 mV. The resistance value measured and averaged over a period of one second is 40 kΩ and is temporarily stored. This determines the ionic conductivity (see S110).

[0222] The voltage-tapping electrodes are then disconnected from the conductivity measuring channel via the measuring device's analog switches, and the counter electrode GE1 5 and reference electrode RE1 6 used for current input are connected together with the working electrode AE1 4 as a potentiostatic three-electrode arrangement in the first measuring chamber 2 and connected to the first voltammetric measuring channel for amperometric measurement of the electrochemically active substances, with the working electrode AE1 4 being subjected to a polarization voltage of 300 mV. In parallel, the second potentiostatic three-electrode arrangement consisting of the working electrode AE2 8, GE 2 9, and RE2 10 is put into operation for the enzymatic amperometric determination of the analyte concentration in the second measuring chamber by connecting the second voltammetric measuring channel, which applies a polarization voltage of 200 mV to the working electrode AE2 8.

[0223] In this example, for example, three seconds after switching on the polarization voltages, the signal currents of both measuring channels are measured and integrated over a period of five seconds, and the raw charge values ​​Q con c = 12.8 pC and Q in tf 1.2 pC are temporarily stored. These charges contain information about the analyte concentration and the interferent concentration. This step corresponds to steps S130 and S120. The electrodes are then disconnected from the measuring channels via the analog switch array in both measuring chambers.

[0224] Finally, a second resistance measurement is performed on the meander trace over a further 500 ms, resulting in a resistance value of 548.4 Ω. The resistance values ​​measured at the beginning and end of 549.1 Ω and 548.4 Ω are averaged, preferably arithmetically. According to equation (1), the averaged resistance value of 548.76 Ω corresponds to a temperature of 15.2 °C according to the calibration curve in Fig. 4. This temperature value is temporarily stored and used for the subsequent evaluation steps.

[0225] The relationship between ionic conductivity or resistance and hematocrit is illustrated by way of example for three different temperatures by the calibration curves in Fig. 5. The resistance value of 40 kΩ, averaged as a function of the ionic conductivity of the first measuring chamber 2 and over an exemplary period of one second, is determined using equation 3 and corresponds to a hematocrit of 45% at 15°C (see step S210 in Fig. 3). The correlation curve or calibration curve between resistance value / ionic conductivity and hematocrit at 15°C, which results from the empirically recorded and pre-stored values ​​for 15°C in Fig. 5, follows equation (16).

[0226] Y = 21.11 + 0.124x + 0.0065x 2(16) In a subsequent step, the signal currents for measuring the analyte concentration and electrochemically active substances must be corrected for the temperature influence, see steps S220 and S230 in Fig. 3. For the charge value of 12.8 pC for lactate, temporarily stored as charge after integration, the intersection point with the nominal temperature of 25°C is determined using the set of curves in Figure 7, which was generated from the set of calibration curves from Figure 6. This point lies between the charge value curves of 10 pC and 15 pC and, when determined graphically, corresponds to a lactate value of approximately 6.5 mM. The two curves adjacent to the determined charge value of 12.8 pC, each with the same charge of 10 pC and 15 pC, run according to equations (17) and (18):

[0227] Qio P c: y = 7.52768 - 0.10833x+0.00035x 2 (17)

[0228] Qi5 M c: y = 13.10747 - 0.27714 x -0.00237 x 2 (18)

[0229] At a nominal temperature of 25°C, according to equations (17), (18) for Q M c = 5.04 mM and for QIS M C: 7.69 mM. After linear iteration, the charge value of 12.8 pC at a nominal temperature of 25°C corresponds to a lactate concentration of 6.52 mM. In an analogous manner and as described in more detail in connection with Figure 3, the concentration value for electrochemically active substances corrected to 25°C is determined to be 0.6 mM.

[0230] The temperature-corrected raw concentration values ​​are now used to correct the hematocrit dependence of the lactate concentration value using the previously recorded and pre-stored calibration curves, as shown in Fig. 8 and Fig. 9, and according to equations (11) to (14), cf. S330 in Fig. 3.

[0231] For the buffered temperature-corrected concentration value of 6.52 mM lactate, the intersection point with the previously determined hematocrit value of 45% is determined using the pre-stored set of curves in Fig. 9, which was generated from the set of calibration curves (charge values ​​vs. lactate concentration) as a function of hematocrit (Fig. 8). In the present example, this lies between the curves of equal charge of 10 pC and 15 pC. The two curves Q10 and Qis from Fig. 9 are described by equations (19) and (20):

[0232] Qio P c: Y = 4, 0649 - 0.02965x 4 - 0.00106 x 2 (19)

[0233] Qi5 M c: Y = 6.07069 - 0.04224x 4 - 0.00164 x 2 (20)

[0234] If x = 0% is used for the hematocrit to obtain the plasma-related value, the plasma lactate values ​​are 4.06 mM and 6.07 mM. Using the lactate values ​​for 45%, a linear iteration is performed, resulting in a hematocrit- and temperature-corrected plasma lactate value of 5.4 mM in this example.

[0235] An analogous procedure was used for the 25°C-corrected interference concentration value of electrochemically active substances. After hematocrit correction, a concentration of 0.4 mM was obtained for the hematocrit-temperature-corrected interference concentration. These steps correspond to the steps S330 and S320 described above in Fig. 3.

[0236] In a further step, the influence of the electrochemically active substances on the determination of the analyte concentration value is corrected by subtraction according to step S430 in Fig. 3. A weighting factor / sensitivity factor Ks according to equation (15) can also be taken into account, which is a quotient of the sensitivity of the amperometric analyte measurement system and the amperometric interference measurement system to electrochemically active substances. Ks is calculated from the quotient of the slope of the straight line equation for MK2 and the slope of the straight line equation for MK1 from Fig. 10 and reflects the influence of electrochemically active substances on the charge values ​​of measuring chamber 1 and 2, respectively. The reason for the different sensitivity of the two measuring chambers to electrochemically active substances is the different reagent composition.Since the reagent mixture for lactate measurement in the second measuring chamber 3 generally has a higher ion concentration and protein content than the reagent system for detecting electrochemically active substances in the first measuring chamber, electrochemically active substances are detected with greater sensitivity in the first measuring chamber due to lower diffusion barriers. In the present case, the sensitivity factor was calculated using K using the calibration curves in Fig. 10. s =0.74, so that in this example the plasma-related lactate value corrected for the influence of electrochemically active interfering substances is 5.1 mM.

[0237] The described method is particularly suitable for single-use sensors that are intended to provide clinically relevant, fast, and accurate plasma-related concentration values. By using a temperature measuring resistor applied to the carrier, i.e., integrated, the current ambient temperature can be measured with high precision and without susceptibility to interference compared to the state of the art (see the above explanations). This precisely measured ambient temperature is then used step by step to systematically correct for interference caused by hematocrit and electrochemically active substances. This allows for systematic elimination of interference, allowing for adequate correction of these interferences. Because the analyte concentration measurement is plasma-related, the sensor is particularly suitable for emergency situations or for rapidly required analyte determinations where whole blood must be used.

[0238] Implementation example 2

[0239] The following figures demonstrate the method described above and the sensor implementing the method using a specific example with D-lactate as the analyte to be determined. For additional steps, reference is made to the procedures described above. Working example 2 further supplements the above information.

[0240] The measurements were carried out as in Example 1, using an NAD-independent D-lactate dehydrogenase from Desulfovibrio vulgaris (Ogata et al. (1981) J. Biochem 89, 1423-1431). The enzyme was obtained from ProteoGenix. It was recombinantly expressed in Escherichia coli and purified by Ni affinity chromatography using an N-terminal His tag.

[0241] According to WO2012 / 01334415, an artificial synovial fluid sample was prepared which contained the following components: 12 mg / mL BSA, 7 mg / mL bovine blood globulins, 0.1 mg / mL dipalmitoylphosphytidylcholine, 2 mg / mL hyaluronic acid, 1 x phosphate buffered saline (1xPBS, pH 7.1).

[0242] These samples were spiked with a defined amount of D-lactate.

[0243] A photometric measurement using a commercial detection kit (D-Lactate Assay Kit, Art. No. MAK336, Sigma-Aldrich) was used as the reference method.

[0244] Figure 11 shows a calibration curve from 0.2 to 2.0 mM D-lactate obtained with these samples in the photometric detection. For the artificial synovial fluid without added D-lactate, a concentration of 0 mM D-lactate was determined. Accordingly, the artificial synovial fluid apparently contained no substances that would interfere with the photometric detection.

[0245] Figure 12 shows measurements of the sensor signal (i.e., charge in pC) with the sensor according to the invention at various amounts of enzyme (lactate dehydrogenase). This demonstrates that the sensor system with potassium hexacyanoferrate(III) as a redox mediator and lactate dehydrogenase is suitable for determining the concentration of D-lactate. The measurements were performed with 2.3 mM FAD (enzyme cofactor) and various amounts of enzyme in 1x PBS, pH 7.4.

[0246] Figure 13 shows the sensor signal as a function of incubation time and incubation temperature. The reagent mixture contained 0 or 10 mM D-lactate, 2.3 mM FAD, and 1200 U D-lactate dehydrogenase in 1x TBS (Tris-buffered saline), pH 8.0, at 37°C. The sensor signal was not significantly affected by temperature (room temperature versus 37°C). This demonstrates that the determination of D-lactate concentration with the sensor according to the invention is possible at room temperature (RT). The signal strength increases with longer incubation times, with a flattening of the curve observed between 30 and 60 minutes.

[0247] Figure 14 shows the sensor behavior under different buffer conditions. For this purpose, the sensor was pre-incubated for 10 minutes in the specified compositions (PP = 50 mM sodium phosphate buffer (pH 6.8), BSA = bovine serum albumin). Slightly alkaline buffer conditions (pH 8.0–8.5) result in higher signal intensities compared to neutral buffer solutions.

[0248] Figure 15 shows the sensor signal for various D-lactate concentrations as a function of incubation time (n=4). The samples were incubated at room temperature for 5 (A), 10 (■), and 20 minutes (•), respectively. A clear dependence of the sensor signal on incubation time was observed. The difference between the signal strengths at different incubation times increased with higher D-lactate concentrations. Based on these data, an incubation time of 10 minutes was determined to be the preferred value, also considering the shortest possible determination time for the sensor user. This value was also used for the following examples.

[0249] Figure 16 shows a bar chart of sensor signals for low D-lactate concentrations (0, 0.5, and 1 mM D-lactate) for 5, 10, and 20 minutes of pre-incubation time. This shows a clear dependence of the sensor signal on the incubation time. The background signal (basic charge) at 0 mM D-lactate also increases slightly with extended incubation time, especially at 20 minutes. Based on these data and the additional data in Figure 15, an incubation time of 10 minutes was determined to be the preferred value, also considering the shortest possible determination time for the sensor user. This value was also used for the following examples. Figure 17 shows the signal-to-noise ratio of the sensor for 10 mM D-lactate and 0 mM D-lactate in the presence of 650 U D-lactate dehydrogenase, 2.3 mM FAD, and 270 mM potassium hexacyanoferrate(III) in 1x PBS (pH 7.4) after 10 minutes of pre-incubation time and at 37°C.The polarization voltage was applied immediately before the measurement. No significant difference was observed between the tested measurement segments. The sensor's signal-to-noise ratio was between 4.4 and 5.1 in all tested segments.

[0250] Figure 18 shows the signal-to-noise performance of the sensor at different polarization voltages for 10 mM D-lactate in the presence of 650 U D-lactate dehydrogenase, 2.3 mM FAD, and 270 mM potassium hexacyanoferrate(III) in 1x PBS after a 10-minute pre-incubation time at 37°C. A polarization voltage of 100 or 200 mV provides the best signal-to-noise performance. A higher polarization voltage could lead to the oxidation of endogenous or exogenous redox-active substances, such as uric acid. Therefore, the use of a lower polarization voltage may be advantageous.

[0251] Figure 19 shows the sensor signal for 10 mM D-lactate (■) and 0 mM D-lactate (•) after switching on the polarization voltage in the 1st second and measuring the charge [pC] during the indicated segments. The measurements were carried out in the presence of 650 U D-lactate dehydrogenase, 2.3 mM FAD, and 270 mM potassium hexacyanoferrate(III) in 1x PBS (pH 7.4) after a 10-minute pre-incubation time at room temperature. The signal-to-noise ratio (A) is also shown. No significant difference was observed between the tested measurement segments. In the following examples, the charge was therefore measured from the seventh to the 10th segment.

[0252] second, and the polarization voltage is switched on immediately before the measuring segment in the 6th second.

[0253] Figure 20 shows dose-response curves for a D-lactate dehydrogenase sensor assay according to the invention. The measurements were carried out with various D-lactate dilutions in buffer (■), single-strength artificial synovial fluid (A), or double-strength artificial synovial fluid (•). The data shown here were obtained according to the measurement conditions optimized in the previous experiments. The error bars correspond to the standard deviation of 2 to 3 sensors. The background value, determined by measuring 0 mM D-lactate, was subtracted from the signal in each case. This background value was approximately 1 pC for buffer, 1.3 pC for single-strength synovial fluid, and 2.1 pC for double-strength artificial synovial fluid.

[0254] Figure 21 shows the determination of the detection limit of a D-lactate sensor according to the invention in buffer (A) and artificial synovial fluid (B). The detection limit is defined as the background value + 3x standard deviation. The detection limit is shown as a dashed line and is 0.04 mM in each case.

[0255] Figure 22 shows the determination of the detection limit in single-concentration artificial synovial fluid. The detection limit is shown as a dashed line (n = 10). The pre-incubation time was 10 minutes at room temperature. Measurements were performed with ten sensors for each D-lactate concentration (including the background value). The detection limit of 0.04 mM D-lactate in artificial synovial fluid was confirmed. At a concentration of 0.04 mM D-lactate, the sensor signal was above the detection limit for all ten of the ten measured sensors. This demonstrates that a reliable detection limit could be determined.

[0256] Figure 23 shows the sensor signal of the D-lactate sensor at the indicated concentrations of various substances, which were tested for interference with the D-lactate measurement. Three different concentrations of D-lactate served as positive controls, while buffer or artificial synovial fluid served as negative controls. The individual potentially interfering substances were each measured in conjunction with buffer (1x TBS, pH 8.5; black bars) or artificial synovial fluid (gray bars). The respective reaction mixture was pre-incubated for 10 minutes at room temperature. This demonstrated that the D-lactate sensor exhibited high specificity for D-lactate and showed no significant interference even at very high concentrations of L-lactate, pyruvate, NAD, and pyruvate, or artificial synovial fluid.However, an interfering effect of uric acid was observed, which was not due to an enzyme reaction but to a direct interaction with the electrode surface.

[0257] Figure 24 shows a dose-response curve of the D-lactate sensor using real patient samples. The measurements were performed with various dilutions of D-lactate in human synovial fluid, with the synovial fluid samples previously pooled (item number 991-42-P, Lee Biosolutions). The samples were used without further preparation. The data shown here were obtained using the optimized measurement conditions outlined above. The error bars represent the standard deviation from 2 to 3 sensors. The background signal from a measurement with 0 mM D-lactate was subtracted from all measured values. The background signal was approximately 3 pC.

[0258] Figure 25 shows the determination of the detection limit for D-lactate using the measurements from Figure 24. A detection limit of 0.075 mM was determined using linear regression. The detection limit is shown as a dashed line. This underscores the robustness of the D-lactate sensor, as the detection limit in real sample material is in a similar range to that in buffer.

[0259] List of reference symbols

[0260] 100 sensors

[0261] 1 carrier

[0262] 2 first measuring chamber

[0263] 3 second measuring chamber

[0264] 4 Working electrode AE1

[0265] 5 Counter electrode

[0266] 6 Reference electrode

[0267] 7a, 7b Voltage tapping electrodes

[0268] 8 Working electrode

[0269] 9 Counter electrode

[0270] 10 Reference electrode

[0271] 11 Temperature measuring resistor / meander conductor structure

[0272] 12 supply lines

[0273] 13 electrical contacts

[0274] 14 Insulating varnish

[0275] 17 Sample recording area

[0276] 18a, b ventilation duct

[0277] D support section

[0278] L total length

[0279] 31 first three-electrode arrangement

[0280] 32 second three-electrode arrangement

[0281] 33 Four-electrode conductivity arrangement

[0282] 50 processor unit

[0283] 55 analog switching array

[0284] 60 memory / data storage

Claims

Patent claims 1. A method for determining an interference-corrected analyte concentration in a sample by a sensor, comprising the steps: Filling at least two measuring chambers (2, 3) with a sample via a sample receiving area (17) of the sensor, wherein the at least two measuring chambers are formed on a support (1) of the sensor and comprise a first three-electrode arrangement (31), a four-electrode conductivity arrangement (33) and a second three-electrode arrangement (32); Determining an ionic conductivity (S110) of the sample with the four-electrode conductivity arrangement; Determining an interference charge (S120) of electrochemically active substances of the sample by the first three-electrode arrangement; enzymatically-electrochemically determining an analyte charge (S130) of the sample by the second three-electrode arrangement; Determining an analyte concentration (230) using pre-stored calibration curves and the determined analyte charge; Determining an interference concentration (S220) using pre-stored calibration curves and the determined interference charge; and determining a contamination value (S210) using pre-stored calibration curves and the determined ionic conductivity; Correcting the analyte concentration and the interference concentration to a contamination-corrected analyte concentration (S330) and a contamination-corrected interference concentration (S320) using pre-stored calibration curves and the previously determined contamination value; Determine the interference-corrected analyte concentration (S430) by subtracting the interference concentration from the analyte concentration, where the analyte is D-lactate.

2. The method according to claim 1, further comprising determining an ambient temperature (S100) after filling the measuring chambers by means of a temperature measuring resistor (11) applied to the carrier.

3. The method according to claim 2, wherein the analyte concentration, the interference concentration and / or the contamination concentration are further determined based on the ambient temperature determined by means of the temperature measuring resistor using pre-stored calibration curves can be corrected.

4. Method according to one of claims 2 to 3, wherein the temperature measuring resistor is a meander conductor structure which is applied to the carrier of the sensor, wherein the meander conductor structure is preferably positioned adjacent to the measuring chambers, further preferably positioned in a carrier section (D) which is less than one third, preferably less than one quarter, even more preferably less than one fifth of the total length (L) of the carrier.

5. The method according to any one of claims 2 to 5, wherein determining the ambient temperature comprises the following steps: Determining a first temperature after filling the measuring chambers; determining a second temperature after determining the ionic conductivity, the analyte charge, and / or the interference charge of electrochemically active substances in the sample; and Determine the ambient temperature by arithmetic averaging the determined temperatures.

6. The method according to any one of claims 1 to 6, wherein the first measuring chamber comprises the first three-electrode arrangement and the four-electrode conductivity arrangement, and the second measuring chamber comprises a second three-electrode arrangement, further comprising switching the electrodes (4, 5, 6, 7a, 7b) in the first measuring chamber between the three-electrode arrangement for electrochemical measurement and the four-electrode conductivity arrangement for measuring ionic conductivity by means of an integrated or reversibly connected analog switch array (55).

7. The method according to any one of claims 1 to 7, wherein the first three-electrode and four-electrode conductivity arrangement has a reagent coating comprising a redox mediator, and wherein the second three-electrode arrangement has a reagent coating comprising a D-lactate dehydrogenase, preferably an NAD-independent D-lactate dehydrogenase, or other catalytically active proteins, and a redox mediator.

8. Method according to one of claims 1 to 7, wherein the sample is treated together with a reagent mixture which comprises a buffer solution, a redox mediator, optionally further additives for stabilization and a D-lactate dehydrogenase, preferably a NAD-independent D-lactate dehydrogenase, is incubated.

9. The method according to claim 8, wherein the sample is incubated with the reagent mixture for 1 to 30 minutes, preferably 10 minutes, at 18 to 28°C, preferably 24°C, before being applied to the sensor.

10. The method according to any one of claims 1 to 9, wherein the determination of the ionic conductivity of the sample by the four-electrode conductivity arrangement, the determination of the interference charge of electrochemically active substances of the sample by the first three-electrode arrangement and the enzymatic-electrochemical determination of the analyte concentration of the sample by the second three-electrode arrangement are carried out stepwise within a measuring interval of 2 s to 60 s, preferably between 8 and 12 s.

11. Sensor for determining an analyte concentration in a sample, comprising: at least two measuring chambers (2, 3) which are connected to a sample via a Sample receiving area (17) of the sensor can be filled, wherein the at least two measuring chambers (2, 3) are formed on a carrier (1) and comprise a first three-electrode arrangement (31), a four-electrode conductivity arrangement (33) and a second three-electrode arrangement (32), at least one processor unit (50) which is reversibly connected to the sensor via an electrical contact (13) or is integrated on the carrier (1), and is configured to do the following: controlling the four-electrode conductivity arrangement (33) to determine an ionic conductivity of the sample; Controlling the first three-electrode arrangement (31) to determine an interference charge of electrochemically active substances of the sample; controlling the second three-electrode arrangement (32) to enzymatically-electrochemically determine an analyte charge of the sample; Determining an analyte concentration (230) using pre-stored calibration curves and the determined analyte charge; Determining an interference concentration (S220) using pre-stored calibration curves and the determined interference charge; and determining a contamination value (S210) using pre-stored calibration curves and the determined ionic conductivity; Correcting the analyte concentration and the interference concentration to a contamination-corrected analyte concentration (S330) and a contamination-corrected interference concentration (S320) using pre-stored calibration curves and the previously determined contamination value; Determine the interference-corrected analyte concentration (S430) by subtracting the interference concentration from the analyte concentration, where the analyte concentration is the concentration of D-lactate in the sample.

12. Sensor according to claim 11, further comprising a temperature measuring resistor (11) applied to the carrier (1) for determining an ambient temperature, wherein the sensor is arranged to control the temperature measuring resistor (11) in order to determine the ambient temperature after filling the measuring chambers (2, 3); 13. Sensor according to claim 12, wherein the temperature measuring resistor (11) is a meander conductor structure which is applied to the carrier (1) of the sensor, wherein the meander conductor structure (11) is preferably positioned adjacent to the measuring chambers (2, 3), more preferably positioned in a carrier section (D) which is less than one third, preferably less than one quarter, even more preferably less than one fifth of the total length (L) of the carrier (1).

14. Sensor according to one of claims 10 to 13, wherein the first measuring chamber (2) comprises the first three-electrode arrangement (31) and the four-electrode conductivity arrangement (33), and the second measuring chamber (3) comprises the second three-electrode arrangement (32), wherein the at least one processor unit (50) is configured to switch electrodes (4, 5, 6, 7a, 7b) in the first measuring chamber (2) between the first three-electrode arrangement (31) for electrochemical measurement and the four-electrode conductivity arrangement (33) for measuring the ionic conductivity by means of an integrated or reversibly connected analog switch array (55).

15. Sensor according to one of claims 10 to 14, wherein the first three-electrode arrangement (31) is a three-electrode arrangement and / or the second three-electrode arrangement (32) is a three-electrode arrangement.

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