Method and sensor for determining a plasma-related analyte concentration in whole blood

Abstract

The invention relates to a method and a sensor for determining a plasma-related analyte concentration in whole blood. An ambient temperature is determined (S100) by means of a temperature-measuring resistor (11) attached to a support (1) of the sensor, said ambient temperature being used to determine (S230, S210, S220) the haematocrit, the interference concentration of electrochemically active substances, and an analyte concentration in the whole blood sample in a temperature-corrected manner. The temperature-corrected haematocrit is then used to determine a plasma-related haematocrit-corrected and temperature-corrected analyte concentration and a plasma-related haematocrit-corrected and temperature-corrected interference concentration (S330, S320). This is followed by determining the analyte concentration (S430) by subtracting the haematocrit-corrected and temperature-corrected interference concentration from the haematocrit-corrected and temperature-corrected analyte concentration.

Description

[0001] Method and sensor for determining a plasma-related analyte concentration in whole blood

[0002] The invention relates to a method and a sensor for determining a plasma-related analyte concentration in whole blood.

[0003] Technological background

[0004] Whole blood measurements of analytes such as lactate or glucose using enzymatic voltammetric 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 the hematocrit of a sample, redox-active endogenous metabolites and medications, and, above all, the ambient temperature. Temperature influences not only the enzymatic indication reaction, but also the viscosity of the sample, 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] When designing disposable sensors with temperature correction, a calibration curve is typically recorded using spiked whole blood 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. For this purpose, however, it is advisable that the current ambient temperature is known or can be measured as accurately as possible and close to the sensor. The most common technical solution for compensating for the temperature influence is the use of a miniaturized PT100 temperature sensor, which is arranged in the handheld measuring device in the immediate vicinity of the plugged-in disposable sensor. Its calibrated temperature measurement is used via an algorithm to correct the measured value.Because the measuring device, unlike a 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 measuring temperature, the blood droplet can lead to a measured 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.

[0006] Other known solutions from the prior art describe the use of an amperometric two-electrode arrangement consisting of a working (anode) and a reference electrode (cathode), 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.

[0007] 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 applicable, the temperature measured by the measuring device in order to achieve compensation for the temperature influence.

[0008] 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.

[0009] 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, allowing both variables to be used to correct the enzymatic voltammetric analyte determination.

[0010] 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.

[0011] Another cause of inaccurate measurements can be the hematocrit of the blood sample. 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 ranges between 40% and 48% (men) and 36% and 42% (women). However, depending on a subject's genetic predisposition, age, gender, health status, and physical activity, values ​​between 20% and 70% can occur.

[0012] Erythrocytes are oval-shaped cells with a diameter of 2 to 30 μm. At 74.6 g / dL (W / V), they have a significantly lower water content 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 that is approximately 10% lower than that measured in plasma.

[0013] 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 lactate, the concentration in the erythrocytes is described as being between 50% and 70% of the plasma value.

[0014] 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.

[0015] However, for hematocrit values ​​outside the normal range, this measurement error caused by the red blood cell volume can increase considerably.

[0016] 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.

[0017] As a result, the measured analyte concentration will be too low for samples with high hematocrit and too high for samples with low hematocrit.

[0018] 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 in order 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 carry out 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 and the second frequency is greater than 20 kHz. The impedance measurement is carried out on a two-electrode arrangement arranged 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.From the two measured values, a multiple regression analysis is carried out 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.

[0019] 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.

[0020] Furthermore, a sensor arrangement with a capillary gap measuring chamber is known, in which two lateral spacer walls are made of conductive material, to which an alternating current measurement is applied 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 each other 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.

[0021] When measuring the conductivity of the sample, however, it must also be taken into account that analytes such as lactic acid or the resulting lactate are present in dissociated form and, with increasing concentration, cause an increase in conductivity, so that the Hct determined via conductivity measurement may be incorrect without taking the analyte concentration into account.

[0022] 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.

[0023] Furthermore, using differential pulse voltammetry 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.

[0024] 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.

[0025] 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.

[0026] 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.

[0027] 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 from the first redox mediator in its formal redox potential by at least 200 mV. 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 (Differential Pulse Voltammetry), 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.

[0028] 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.

[0029] 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 HOT, 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.

[0030] 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.

[0031] Another approach using this amperometric electrode arrangement aims to measure a so-called test current value immediately after the polarization voltage is applied and after 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.

[0032] 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.

[0033] 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.

[0034] 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.

[0035] 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.

[0036] 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.

[0037] In a 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 lactate measurement. 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.

[0038] 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.

[0039] A known technical solution uses disposable sensors for the determination of glucose, lactate and creatinine in a capillary gap arrangement with an additional working electrode that 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.

[0040] 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.

[0041] 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.

[0042] 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.

[0043] 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 a plasma-related analyte concentration in whole blood. This enables a more accurate enzymatic-voltammetric analyte determination during the measurement procedure, compensating for the actual temperature, the hematocrit of the sample, and the electrochemically active substances contained therein, so that these interferences on the plasma-related analyte measurement value to be determined can be adequately corrected. Further problems solved will become apparent from the following description.

[0044] The object of the invention is achieved according to the independent claims. The subclaims represent preferred embodiments.

[0045] Summary of the invention

[0046] A first aspect of the invention relates to a method for determining an analyte concentration in whole blood using a sensor. In one step, at least two measuring chambers are filled with a whole blood 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 voltammetric three-electrode arrangement, a four-electrode conductivity arrangement, and a second voltammetric three-electrode arrangement. In one 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 whole blood sample using the four-electrode conductivity arrangement.The method further comprises the step of voltammetrically determining an interference charge of electrochemically active substances in the whole blood sample using the first voltammetric three-electrode arrangement. The method further comprises the step of enzymatically voltammetrically determining an analyte charge in the whole blood sample using the second voltammetric three-electrode arrangement. Furthermore, the method comprises determining a temperature-corrected analyte concentration using pre-stored calibration curves, the determined ambient temperature, and the determined analyte charge.The method further comprises determining a temperature-corrected interference concentration using pre-stored calibration curves, the determined ambient temperature, and the determined interference charge; and further comprising determining a temperature-corrected hematocrit value using pre-stored calibration curves, the determined ambient temperature, and the determined ionic conductivity. The method further comprises correcting the temperature-corrected analyte concentration and the temperature-corrected interference concentration to a plasma-related hematocrit- and temperature-corrected analyte concentration and a plasma-related hematocrit- and temperature-corrected interference concentration using pre-stored calibration curves and the previously determined temperature-corrected hematocrit value, respectively.The method further comprises the step of determining the analyte concentration by subtracting the hematocrit and temperature corrected interference concentration from the hematocrit and temperature corrected analyte concentration.

[0047] The analyte can be, for example, glucose, lactate, or creatinine, although the invention is not limited thereto. Electrochemically active substances can be, for example, ascorbate, uric acid, bilirubin, or paracetamol. Plasma-related means that the analyte concentration is related to the value of a corresponding plasma sample (hematocrit concentration = 0%). Potentiostatic three-electrode arrangements are preferably used as indication systems for the two voltammetric measurements. During subtraction, a sensitivity factor, set in Example 1, can preferably be determined, which weights the hematocrit- and 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 nonspecific.The enzymatic voltammetric measurement provides the selective or analyte-specific measurement signal, which, however, has a non-specific component due to the voltammetric detection principle and is corrected by the simple voltammetric measurement.

[0048] For temperature correction, the calibration curves can include corresponding pre-stored, i.e., previously recorded, calibration curves between the analyte concentration and temperature, between the interference concentration and temperature, and between hematocrit / ionic conductivity and temperature. For hematocrit correction, i.e., plasma-related correction, hematocrit is used using previously determined calibration curves / sets of correlation curves between analyte concentration and hematocrit and between interference concentration and hematocrit. The pre-stored calibration curves can be stored in a memory / data storage device, which can also be accessed by a processor implementing the method.

[0049] The first measuring chamber and the second measuring chamber can each preferably have a volume of 0.15 pL to 0.3 pL. The temporal sequence can be as follows: (i) The temperature measurement can be carried out immediately after filling the measuring chambers and after temporally defined current thresholds of the voltammetric measuring channels have been exceeded over 50 to 1000 ms at the resistor, (ii) An ionic conductivity measurement for hematocrit determination can follow at a time interval of 0.5 s to 1 s, (iii) The voltammetric measurement for determining the interference concentration in the whole blood sample can be carried out between the second and ninth second, (iv) An enzymatic voltammetric measurement of the analyte concentration can be carried out between the third and tenth second. Thus, the entire measurement process is completed in approximately 10 s, at least less than 20 s.

[0050] The described method is particularly suitable for disposable sensors intended to provide clinically relevant and reliable plasma-related concentration values. A further technical advantage 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 state of the art (see the above explanations). This precisely measured ambient temperature is then also used to correct the temperature-dependent enzymatic voltammetric analyte measurement and to eliminate the temperature-dependent interference caused by hematocrit and electrochemically active substances. The influence of hematocrit and interfering electrochemically active substances is thus corrected more systematically and adequately than with known technical solutions.Because the analyte concentration is determined based on plasma, the disposable sensor is particularly suitable for emergency situations or for rapidly required analyte determinations. In contrast to clinical settings, where the hematocrit can be removed from the whole blood prior to automated analyte measurements or where blood gas analyzers are used, which require significantly larger sample volumes and greater time and equipment expenditure to prepare the hematocrit, a very small whole blood sample is used in these application areas, allowing the analyte value to be measured in seconds, even under field conditions, in a clinically relevant and reliable manner.

[0051] The method can preferably further comprise correcting the temperature-corrected hematocrit value to an analyte- and temperature-corrected hematocrit value 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. Thus, a correction of the ionic conductivity measurement can be made due to ionically conductive components of the sample, in particular a dissociated analyte or dissociated interference concentrations, such as endogenous or exogenous metabolites. The hematocrit value is thus determined more accurately, so that the analyte concentration to be determined, which depends on this determined hematocrit value, and the interference concentration to be deducted can also be determined more accurately.The calibration curves must be determined between ionic conductivity / hematocrit and analyte concentration or between ionic conductivity / hematocrit and interference concentration. The temperature- and analyte-corrected conductivity measurement can be determined using a previously determined, nonlinear calibration curve between hematocrit and ionic conductivity of the sample to determine the hematocrit value.

[0052] The 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 plasma-related analyte concentration to be determined with greater precision. The meander conductor structure preferably has a resistance between 100 O and 2000 O and a temperature coefficient between 0.4 O / °C and 0.7 O / °C for temperature-dependent resistance measurement.

[0053] 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.

[0054] 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 plasma-related analyte concentration to be determined can also be determined with greater accuracy.

[0055] Preferably, determining the ambient temperature comprises determining a first temperature after or during filling the measuring chambers; determining a second temperature after determining the hematocrit value, the analyte charge, and / or the interference charge of electrochemically active substances in the whole blood 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 plasma-related analyte concentration to be determined can also be determined with greater accuracy, since the undesired temperature drift is compensated or averaged out. The two temperature-dependent resistance measurements on the meander conductor structure can be performed over a period of 50 to 1000 ms.

[0056] The first measuring chamber preferably comprises the first voltammetric three-electrode arrangement and the four-electrode conductivity arrangement; the second measuring chamber preferably comprises the second voltammetric three-electrode arrangement. 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 meandering conductor structure.This in turn improves the accuracy of temperature compensation, so that the plasma-related analyte concentration to be determined can also be determined with greater accuracy.

[0057] 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.

[0058] 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.

[0059] 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.

[0060] 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.

[0061] Preferably, the method comprises measuring the ionic conductivity of the whole blood sample by the four-electrode conductivity arrangement, voltammetrically determining the interference charge of electrochemically active substances of the whole blood sample by the first voltammetric three-electrode arrangement, and enzymatically voltammetrically determining the analyte charge of the whole blood sample by the second voltammetric three-electrode arrangement, step by step within a measuring interval of 8 s to 20 s, preferably between 8 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 for temperature determination can be carried out over 50 ms to 1000 ms, (ii) An ionic conductivity measurement for hematocrit determination can follow with a time interval of 0.5 s to 1 s, (iii) The voltammetric measurement for determining the interference concentration in the whole blood sample in the first measuring chamber can be carried out between the second and ninth second, (iv) An enzymatic voltammetric measurement of the analyte concentration can be carried out between the third and tenth second, (v) A second measurement of the meander resistance for temperature determination can be carried out over 50 ms to 1000 ms.

[0062] In a further aspect of the invention, a sensor for determining a plasma-related analyte concentration in whole blood is described. This sensor comprises at least two measuring chambers that can be filled with a whole blood sample via a sample receiving area of ​​the sensor, wherein at least two measuring chambers are formed on a carrier and comprise a first voltammetric three-electrode arrangement, a four-electrode conductivity arrangement, and a second voltammetric three-electrode arrangement. The sensor further comprises 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: control the temperature measuring resistor in order to measure the ambient temperature after the measuring chambers have been filled.The functionality further includes the step of controlling the four-electrode conductivity arrangement to determine an ionic conductivity of the whole blood sample. Furthermore, the step of controlling the first three-electrode arrangement to voltammetrically determine an interference charge of electrochemically active substances in the whole blood sample. Furthermore, the step of controlling the second three-electrode arrangement to enzymatically voltammetrically determine an analyte charge in the whole blood sample. Furthermore, the step of determining a temperature-corrected analyte concentration using pre-stored calibration curves, the determined ambient temperature, and the determined analyte charge is included. Furthermore, the step of determining a temperature-corrected interference concentration using pre-stored calibration curves, the determined ambient temperature, and the determined interference charge is included.

[0063] Furthermore, the method includes determining a temperature-corrected hematocrit value using pre-stored calibration curves, the determined ambient temperature, and the determined ionic conductivity. A further step involves correcting the temperature-corrected analyte concentration and the temperature-corrected interference concentration to a plasma-related hematocrit- and temperature-corrected analyte concentration and a plasma-related hematocrit- and temperature-corrected interference concentration using pre-stored calibration curves and the previously determined temperature-corrected hematocrit value. Furthermore, the analyte concentration is determined by subtracting the hematocrit- and temperature-corrected interference concentration from the hematocrit- and temperature-corrected analyte concentration. In other words, the processor unit is a processor or a microcontroller.The pre-stored calibration curves can be stored in a memory / data storage to which the processor unit is operatively connected.

[0064] This results in the same advantages as described for the above procedure, to which reference is hereby made.

[0065] 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.

[0066] The meander ladder structure is 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.

[0067] Preferably, the first measuring chamber comprises the first voltammetric 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 is to be carried out by means of an integrated or reversibly connected analog switch array.

[0068] Further preferred embodiments can also be taken from the above method.

[0069] Short description of the characters

[0070] The invention is explained in more detail below using an exemplary embodiment and the accompanying drawings. The figures show:

[0071] Fig. 1 shows a sensor according to the invention according to a preferred embodiment of the invention;

[0072] Fig. 2 is a schematic representation of a sensor according to a preferred embodiment of the invention;

[0073] Fig. 3 is a schematic representation of a flow chart according to the invention;

[0074] Fig. 4 shows an exemplary pre-stored calibration curve of the meander resistance measurement value as a function of the ambient temperature;

[0075] Fig. 5 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;

[0076] Fig. 6 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;

[0077] Fig. 7 shows an exemplary pre-stored set of curves generated from calibration curves of Fig. 6 by plotting the lactate concentration against the temperature for different charge values ​​measured in the second measuring chamber;

[0078] Fig. 8 exemplary pre-stored calibration curves for charge values ​​as a function of the lactate concentration at hematocrit concentrations of 0% (plasma), 19%, 45% and 70% measured in the second measuring chamber;

[0079] Fig. 9 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, and

[0080] Fig. 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.

[0081] Detailed description of the invention

[0082] 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.

[0083] 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.

[0084] 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 areas of the associated electrode arrangements. Furthermore, a meandering conductor structure of a 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.

[0085] 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

[0086] 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.

[0087] 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.

[0088] 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.

[0089] 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

[0090] 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.

[0091] 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, 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.

[0092] 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.

[0093] The measuring chambers 2, 3 comprise a first voltammetric three-electrode arrangement 31, a four-electrode conductivity arrangement 33 and a second voltammetric three-electrode arrangement 32. These are described in more detail below using a preferred embodiment.

[0094] 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.

[0095] The electrodes 4, 5, 6, 7a, 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. The first voltammetric three-electrode arrangement 4, 5, 6 in the first measuring chamber 2 determines the interference concentration of electrochemically active substances in the whole blood sample. The four-electrode conductivity measuring arrangement 5, 6, 7a, b is used in the first measuring chamber 2 to determine the hematocrit value of the whole blood sample based on ionic conductivity.

[0096] The second voltammetric three-electrode arrangement 32 in the second measuring chamber 3 further comprises a second voltammetric 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 whole blood sample using enzymatic volatometric methods. 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.

[0097] 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 realizing a four-wire resistance measurement. The meander conductor structure 11 and supply lines 13 can be covered with the insulating varnish layer 14. The meander conductor structure 11 is preferably also left uncoated by the insulating varnish layer 14. The meander resistance can be between 100 Ω and 2000 Ω at 25°C.

[0098] 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.

[0099] The pre-stored calibration curves can be stored in a memory / data storage device 60, to which the at least one processor 50 is operatively connected. 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 the 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, 250 mV, and the polarization voltage for the amperometric measurement of the electrochemically active substances is +300 mV, respectively, against the internal reference electrodes RE1 and RE2 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.

[0100] 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.

[0101] Fig. 3 shows a schematic representation of a flow chart according to the invention. The determination of a plasma-related analyte value, for example, for lactate, glucose, etc., which is corrected for the interfering influences of 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.

[0102] In a first step, not explicitly shown here, the at least two measuring chambers 2, 3 (see Fig. 1) are filled with a whole blood sample / capillary blood sample 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 voltammetric three-electrode arrangement 31, a four-electrode conductivity arrangement 33, and a second voltammetric three-electrode arrangement 32.

[0103] 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.

[0104] 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.

[0105] T — kyemp Rvä "*■ väO (1 )

[0106] Here, kyemp - the temperature coefficient (temperature sensitivity) of the meander conductor

[0107] R Mä - the measured meander resistance

[0108] Rvao - the meander resistance at 0°C

[0109] The temperature coefficient (k Te mp) was previously experimentally measured over a range of 0 D C and 50°C recorded calibration curve as a function of the resistance value of the meander conductor T = f (R Mä) 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.

[0110] In a further step, the ionic conductivity S110 of the whole blood sample is determined using the four-electrode conductivity arrangement 33.

[0111] Furthermore, the temperature-corrected hematocrit value 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.

[0112] 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.

[0113] 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.

[0114] 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):

[0115] Rnct = fr (Hct) = ki * concHet 2 + k2* concHet + k3T E {1°C, 5°... 45°C}

[0116] (2)

[0117] Here, the following means:

[0118] 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 blood sample) ki and k2- cell constants fT (Hct) - temperature-dependent parabolic family

[0119] A required semi-parabolic family f T The hematocrit (Hct) is determined using blood samples with hematocrit values ​​between, for example, 0% and 70% for temperatures between preferably 1°C and 45°C in at least 7 steps. The previously determined ambient temperature is then used to assign the hematocrit-dependent resistance measurement value or hematocrit value in the parabolic family.

[0120] The calculation of the temperature-corrected hematocrit can be performed by rearranging equation (2) as follows: concHet

[0121] For resistance values ​​that lie between two isothermal hematocrit semiparabolas of the array, a linear interpolation of the hematocrit value is performed based on the measured temperature. This results in a temperature-corrected hematocrit value.

[0122] In a further step of the method, the voltammetric determination of an interference concentration S120 of electrochemically active substances in the whole blood sample is carried out by the first voltammetric 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 voltammetric 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.

[0123] In a further step, the enzymatic-voltammetric determination of an analyte concentration or analyte charge S130 of the whole blood sample is carried out by the second voltammetric 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-voltammetric determination of the analyte concentration. The provided polarization voltage is preferably between 100 mV and 500 mV.

[0124] The evaluation can be carried out in the following preferred manner: The measured signal currents for measuring the analyte concentration and electrochemically active substances are essentially described by the Cotrell equation (4):

[0125] Herein mean:

[0126] I - current z - number of transferred electrons F - Faraday constant (96.485 As / mol) D - diffusion constant

[0127] A - Electrode surface of the working electrode t - Time c - Initial concentration of the reacted substance

[0128] 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 are also dependent on the temperature and the hematocrit.

[0129] 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.

[0130] For this purpose, previously recorded sets of curves are used according to the following function sets and and which, depending on the analyte or interferential substance concentration of the blood samples, are recorded for example for temperatures between 1°C and 45°C in at least seven temperature levels at an average hematocrit concentration.

[0131] Here, the following means: - temperature-dependent analyte charge values ​​of the curve sets - temperature-dependent interference charge values ​​of the curve families

[0132] T - Ambient temperature (curve parameter) conc Anaiyt - Analyte concentration conc mterf. - Interfering substance concentration kAi , kA2 - Analyte sensitivity k A3 - Charge value at cAnaiyt = 0 mM kn, ki2 — Interference substance sensitivity ki3- Charge value at Cmtf = 0 mM

[0133] The function groups (5) and (6) can then be switched according to the concentration and the concentration values ​​can be calculated using a series of closely spaced

[0134] Charge values ​​(n) are plotted against temperature, resulting in a set of n curves, each of whose functions is described by a general quadratic equation (7), (8): cone Analyt-Tcorr = f QTA (T) = kAT1 T 2 +kAT2 T + kAT3 QTA E {0, 5 C, 10 pC ... 60 pC} (7)

[0135] COnC Interf-Tcorr — fQTI (T) — k|T1 T 2 +kiT2 T + kiT3 QTI G {0, 0.3 pC, 0.6 pC ... 6.0 pC} (8)

[0136] Where: conc mterf-Tcorr- temperature-corrected analyte value conc mterf-Tcorr- temperature-corrected interferent value

[0137] QTA - temperature-dependent analyte charge values ​​(curve family parameters)

[0138] QTI - temperature-dependent interference charge values ​​(curve family parameters)

[0139] T - Ambient temperature kATi, k A T2 - Temperature sensitivity of the analyte measurement kAT3 - Analyte concentration value at T = 0°C kiTi , kiT2 - Temperature sensitivity of the interference measurement kiT3 - Interference concentration value at T = 0°C

[0140] 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.

[0141] 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.

[0142] 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.

[0143] In a preferred embodiment, the correction of the temperature-corrected hematocrit value to an analyte- and temperature-corrected hematocrit value (S240) can be carried out 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.

[0144] This is described in more detail below using an example. The temperature-corrected raw values ​​of the analyte and interference concentrations are used to correct the hematocrit value according to Equation (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.

[0145] 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:

[0146] RHC125°C ( — 1 / GHC — f (Cdis) - adis c dis bdis (9)

[0147] Here, the following means:

[0148] RHCI 25°C - hematocrit-dependent resistance value, determined from the ionic conductivity Gnct in the first measuring cell at 25°C

[0149] Cdis- concentration of a dissociated substance concentration

[0150] -adis - negative increase of the resistance Rnct bdis - resistance Rnct in the absence of a dissociated substance concentration

[0151] 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):

[0152] Rldct-corr = Bdis * COHCRW + RHct 25°C ( )

[0153] Here, the following means:

[0154] COHCRW - raw value of the concentration of a dissociated substance concentration adis - increase in resistance Rnct25°c at 25°C as a function of the concentration of the dissociated substance concentration Rnct-corr- hematocrit-dependent resistance measurement value, corrected for the reduction in the resistance value due to the presence of a dissociated substance concentration at 25°C.

[0155] In a further step, the temperature-corrected analyte concentration and the temperature-corrected interference concentration are now 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.

[0156] 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:

[0157] Q Hct Analyt — fldct (COHC analyte) — SA COHC Analyt Temp corr + bA Hct e {0%, 20%...70% (11) and

[0158] Q Hct Interf — fhct (COHC Interf) — aint COHC Interf Temp corr + blnt Hct e {0%,

[0159] 20%...70%} (12)

[0160] Here, the following means:

[0161] Hct - Hematocrit value (line parameter of the temperature-corrected analyte concentration curves and interferent concentration curves)

[0162] Qnct-Anaiyt - hematocrit-dependent charge values ​​of the temperature-corrected analyte concentration values

[0163] Q Hct-mterf - hematocrit-dependent charge values ​​of the temperature-corrected analyte concentration values ​​of the interferent concentration values ​​cone Anaiyt Temp corr - temperature-corrected analyte concentration values ​​cone interf Temp com — Temperature-corrected interference concentration values ​​a Increase (analyte sensitivity) b - Constant, charge value at cAnaiyt = 0 mM a mt- Increase (interference substance sensitivity) b int - Constant, charge value at cInterfer = 0 mM

[0164] Based on the functional equations (11), (12), the respective concentration values ​​can be determined from the measured charge values ​​Qnct-Anaiyt and QHct-mterf 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): pC, 10 pC....60 pC} (13)

[0165] COnC|nterf-T+Hct corr — fQTHct-l (Hct) - k-THct 11 Hct 2 +k THct I2 Hct + k THct I3 QTHct- I 6 (OpC,

[0166] 0.3 pC, 0.6 pC....6.0 pC} (14)

[0167] Where: conc Anaiyt-T+Hctcorr - temperature and hematocrit corrected analyte value conc interf-T+Hct com - temperature and hematocrit corrected interferent value

[0168] Q THct-A - Temperature-compensated hematocrit-dependent analyte charge (curve family parameter)

[0169] Q THct-l - Temperature-compensated hematocrit-dependent interference discharge of the (curve family parameters)

[0170] Hct - temperature-corrected hematocrit value k T Hcti Ai , k T Hct A2 - temperature-compensated hematocrit sensitivity of the analyte measurement k THct 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)

[0171] 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.

[0172] 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.

[0173] 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).

[0174] 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.

[0175] 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):

[0176] COnC Analyt_T+Hct+lnt -corr — COHC Analyt_T+Hct-corr C int_T+Hct corr * K S (15)

[0177] Where: conc Anaiyt_T+Hct+int -corr plasma-related analyte concentration, corrected for the influence of ambient temperature, hematocrit and electrochemically active substances conc Anaiyt_T+Hct-corr plasma-related analyte concentration, corrected for the influence of ambient temperature and hematocrit

[0178] C int_T+Hct corr Plasma-related total concentration of electrochemically active substances, corrected for the influence of ambient temperature and hematocrit

[0179] K s Sensitivity factor (quotient of sensitivity of the amperometric analyte measuring system and the amperometric interference measuring system against electrochemically active substances). In special versions, K s =1 if a similar reagent composition is assumed in both measuring chambers.

[0180] 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.

[0181] Implementation example

[0182] In the following Figures 4 to 10, the method described above and the sensor implementing the method are demonstrated using a concrete example with 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.

[0183] 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 consisting of redox mediator (20 pg / mL), sodium chloride (2 mM), Tergitol (0.3% VA / ), 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.

[0184] For lactate, the electrodes 8, 9, 10 within the second measuring chamber 3 are preferably coated with a reagent solution containing lactate oxidase (2 pg / mL), sodium chloride (50 mM), CMC (0.5% W / V), 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 exemplary, 560 ohms at 25°C and has a temperature dependence of 0.6 Ω / °C; see Fig. 4.

[0185] 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.

[0186] Immediately after filling the measuring chambers with a capillary blood sample and exceeding temporally 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.

[0187] 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).

[0188] 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.

[0189] 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 CO nc = 12.8 pC and Qmtf = 1.2 pC. 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 in both measuring chambers via the analog switch array.

[0190] Finally, a second resistance measurement is performed on the meander trace over a further 500 ms, yielding a resistance value of 548.4 Ω. The resistance values ​​measured at the beginning and end, 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.

[0191] 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 kO, 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).

[0192] Y = 21.11 + 0.124x + 0.0065x 2 (16)

[0193] In a subsequent step, the signal currents for measuring the analyte concentration and electrochemically active substances must be corrected for temperature influences, 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 in 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):

[0194] QIO M C: y = 7.52768 - 0.10833x+0.00035x 2 (17) QI5 M C: y = 13.10747 - 0.27714x+0.00237x 2 (18)

[0195] At a nominal temperature of 25°C, equations (17) and (18) yield QIO C = 5.04 mM and QI5 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. Analogously, 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.

[0196] 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.

[0197] 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 curve set 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 Qi0 and Qis from Fig. 9 are described by equations (19) and (20):

[0198] QC: Y = 4.0649 - 0.02965x + 0.00106x 2 (19)

[0199] Qi5 M c: Y = 6.07069 - 0.04224x + 0.00164x 2 (20)

[0200] 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.

[0201] In an analogous manner, the 25 D C corrected interference concentration value of electrochemically active substances and, after hematocrit correction, obtain a concentration of 0.4 mM for the hematocrit-temperature corrected interference concentration. These steps correspond to the steps S330 and S320 described above in Fig. 3.

[0202] 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.

[0203] 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 use or for rapidly required analyte determinations that require whole blood.

[0204] 100 sensors

[0205] 1 carrier

[0206] 2 first measuring chamber

[0207] 3 second measuring chamber

[0208] 4 Working electrode AE1

[0209] 5 Counter electrode

[0210] 6 Reference electrode

[0211] 7a, 7b Voltage tapping electrodes

[0212] 8 Working electrode

[0213] 9 Counter electrode

[0214] 10 Reference electrode

[0215] 11 Temperature measuring resistor / meander conductor structure

[0216] 12 supply lines

[0217] 13 electrical contacts

[0218] 14 Insulating varnish

[0219] 17 Sample recording area

[0220] 18a, b ventilation duct

[0221] D support section

[0222] L total length

[0223] 31 first voltammetric three-electrode arrangement

[0224] 32 second voltammetric three-electrode arrangement

[0225] 33 Four-electrode conductivity arrangement

[0226] 50 processor unit

[0227] 55 analog switching array

[0228] 60 memory / data storage

Claims

Patent claims 1. Method for determining a plasma-related analyte concentration in whole blood by a sensor, comprising the steps: Filling at least two measuring chambers (2, 3) with a whole blood sample via a sample receiving area (17) of the sensor, wherein the at least two measuring chambers are formed on a carrier (1) of the sensor and comprise a first voltammetric three-electrode arrangement (31), a four-electrode conductivity arrangement (33) and a second voltammetric three-electrode arrangement (32); Determining an ambient temperature (S100) after filling the measuring chambers by means of a temperature measuring resistor (11) applied to the carrier; Determination of the ionic conductivity (S110) of the whole blood sample using the four-electrode conductivity setup; voltammetric determination of the interference charge (S120) of electrochemically active substances in the whole blood sample using the first voltammetric three-electrode setup; enzymatic-voltammetric determination of the analyte charge (S130) of the whole blood sample using the second voltammetric three-electrode setup; Determining a temperature-corrected analyte concentration (230) using pre-stored calibration curves, the determined ambient temperature and the determined analyte load; Determining a temperature-corrected interference concentration (S220) using pre-stored calibration curves, the determined ambient temperature, and the determined interference charge; and determining a temperature-corrected hematocrit value (S210) using pre-stored calibration curves, the determined ambient temperature, and the determined ionic conductivity; Correcting the temperature-corrected analyte concentration and the temperature-corrected interference concentration 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; Determining the analyte concentration (S430) by measuring the hematocrit and The temperature-corrected analyte concentration is subtracted from the hematocrit and temperature-corrected interference concentration.

2. The method of claim 1, further comprising correcting the temperature-corrected hematocrit value to an analyte- and temperature-corrected hematocrit value (S240) 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.

3. Method according to one of claims 1 to 2, wherein the temperature measuring resistor is a meandering conductor structure which is applied to the carrier of the sensor.

4. Method according to claim 3, wherein the meandering ladder structure is positioned adjoining the measuring chambers.

5. Method according to one of claims 3 to 4, wherein the meandering ladder structure is positioned in a support section (D) which is less than one third, preferably less than one quarter, more preferably less than one fifth of the total length (L) of the support.

6. A method according to any one of claims 1 to 5, wherein determining the ambient temperature comprises the following steps: Determining an initial temperature after filling the measuring chambers; Determining a second temperature after determining the ionic conductivity, analyte charge, and / or interference charge of electrochemically active substances in the whole blood sample; and Determining the ambient temperature by arithmetic averaging the determined temperatures.

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

8. Method according to any one of claims 1 to 7, wherein the first three-electrode and four-electrode conductivity arrangement comprises a reagent coating comprising a redox mediator, and wherein the second three-electrode arrangement comprises a reagent coating comprising an oxidoreductase or further catalytically active proteins and a redox mediator.

9. Method according to any one of claims 1 to 8, wherein the determination of the ionic conductivity of the whole blood sample by the four-electrode conductivity arrangement, the voltammetric determination of the interference charge of electrochemically active substances of the whole blood sample by the first voltammetric three-electrode arrangement and the enzymatic-voltammetric determination of the analyte concentration of the whole blood sample by the second voltammetric three-electrode arrangement are carried out stepwise within a measurement interval of 8 s to 20 s, preferably between 8 and 11 s.

10. Sensor for determining a plasma-related analyte concentration in whole blood, comprising: at least two measuring chambers (2, 3) which can be filled with a whole blood sample via a sample receiving area (17) of the sensor, wherein the at least two measuring chambers (2, 3) are formed on a carrier (1) and comprise a first voltammetric three-electrode arrangement (31), a four-electrode conductivity arrangement (33) and a second voltammetric three-electrode arrangement (32), a temperature measuring resistor (11) applied to the carrier (1) for determining an ambient temperature; 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 perform the following: Controlling the temperature measuring resistor (11) to determine the ambient temperature after filling the measuring chambers (2, 3); Controlling the four-electrode conductivity arrangement (33) to achieve a to determine the ionic conductivity of the whole blood sample; Control of the first three-electrode arrangement (31) to determine an interference charge of electrochemically active substances of the whole blood sample by voltammetry; Control of the second three-electrode arrangement (32) to determine an analyte charge of the whole blood sample enzymatically-voltammetrically; Determining a temperature-corrected analyte concentration using pre-stored calibration curves, the determined ambient temperature, and the determined analyte charge; determining a temperature-corrected interference concentration using pre-stored calibration curves, the determined ambient temperature, and the determined interference charge; and determining a temperature-corrected hematocrit value using pre-stored calibration curves, the determined ambient temperature, and the determined ionic conductivity; Correcting the temperature-corrected analyte concentration and the temperature-corrected interference concentration to a plasma-related hematocrit and temperature-corrected analyte concentration and a plasma-related hematocrit and temperature-corrected interference concentration using pre-stored calibration curves and the previously determined temperature-corrected hematocrit value; Determining the analyte concentration by subtracting the hematocrit- and temperature-corrected interference concentration from the hematocrit- and temperature-corrected analyte concentration.

11. Sensor according to claim 10, wherein the temperature measuring resistor (11 ) is a meandering conductor structure which is applied to the carrier (1) of the sensor.

12. Sensor according to claim 11, wherein the meandering conductor structure (11) is positioned adjoining the measuring chambers (2, 3).

13. Sensor according to one of claims 11 to 12, wherein the meandering conductor structure (11 ) is positioned in a carrier section (D) which is less than one third, preferably less than one quarter, 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) The first measuring chamber (2) comprises a first voltammetric three-electrode arrangement (31) and a four-electrode conductivity arrangement (33), and the second measuring chamber (3) comprises the second voltammetric three-electrode arrangement (32), wherein the at least one processor unit (50) is configured to perform switching of electrodes (4, 5, 6, 7a, 7b) in the first measuring chamber (2) between the first three-electrode arrangement (31) for voltammetric measurement and the four-electrode conductivity arrangement (33) for measuring ionic conductivity by means of an integrated or reversibly connected analog switch array (55).

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