Abstract
A method for detecting in-vivo properties of a biosensor. In the inventive method, a sensitivity-to-admittance relation is provided and a raw current in the biosensor is measured. An in-vivo current response is also measured at first and second operating points. A time constant τ is determined by the electrical capacitance C of the working electrode and the electrical resistance R.sub.M of the membrane by τ=R.sub.M.Math.C. The first and second operating points are selected below and above τ, respectively. An analyte value in a sample of a body fluid is determined by using the raw current and compensating sensitivity drift in the biosensor, which in turn is compensated by using the measured value for the raw current and a corrected value for the sensitivity. The failsafe operation of the biosensor is monitored by using the in-vivo current response measured at the first and second operating points.
Claims
1. A method for detecting in-vivo properties of a biosensor adapted for determining a value of an analyte in a body fluid sample, the biosensor having a working electrode covered by a membrane and an enzyme for providing a reaction with the analyte, the method comprising: a) providing a sensitivity-to-admittance relation of the biosensor; b) measuring a raw current in the biosensor; c) measuring an in-vivo current response indicative of the in-vivo admittance of the biosensor, wherein the in-vivo current response is measured at first and second operating points and a time constant τ is determined by the electrical capacitance C of the working electrode and the electrical resistance R.sub.M of the membrane by τ=R.sub.M.Math.C, wherein the first operating point is selected below τ and the second operating point is selected above τ; d) determining an analyte value in a sample of a body fluid by using the raw current and compensating an in-vivo sensitivity drift in the biosensor, wherein the in-vivo sensitivity drift is compensated by using the measured value for the raw current and a corrected value for the sensitivity, whereby the sensitivity is determined by using the sensitivity-to-admittance relation from step a); and e) monitoring a failsafe operation of the biosensor by using the in-vivo current response measured at the first and second operating points.
2. The method according to claim 1, wherein the first operating point is selected for providing a first characteristic value related to the electrical resistance of the membrane and wherein the second operating point is selected for providing a second characteristic value related to the electrical capacitance of the working electrode.
3. The method according to claim 2, wherein the in-vivo current response of the biosensor is determined by applying a potential step to an electrical potential difference at the biosensor, wherein the potential step comprises applying an additional electrical potential between the working electrode and a reference electrode of the biosensor over a time interval.
4. The method according to claim 1, wherein the second operating point is selected above 3τ.
5. The method according to claim 4, wherein the second operating point is selected above 5τ.
6. The method according to claim 1, wherein the failsafe operation of the biosensor is monitored by using at least one of (i) the sensitivity determined from the sensitivity-to-admittance relation of the biosensor, (ii) the electrical capacitance C of the working electrode, and (iii) the electrical resistance R.sub.M of the membrane.
7. The method according to claim 6, wherein a structural modification of the biosensor is determined by monitoring alterations of at least two of (i) the sensitivity determined from the sensitivity-to-admittance relation, (ii) the electrical capacitance C of the working electrode, and (iii) the electrical resistance R.sub.M of the membrane.
8. The method according to claim 1, wherein the sensitivity-to-admittance relation is obtained during a calibration of the biosensor, wherein the calibration of the biosensor is selected from at least one of a multiple calibration, an initial calibration, and a factory calibration.
9. The method according to claim 1, wherein the biosensor is a fully or partially implantable biosensor for continuously monitoring the analyte.
10. The method according to claim 1, wherein the analyte comprises glucose and the analyte value is determined by using glucose oxidase or glucose dehydrogenase as the enzyme.
11. An electronics unit for detecting in-vivo properties of a biosensor by performing a method according to claim 1.
12. The electronics unit according to claim 11, wherein the electronics unit is further adapted for: applying an electrical potential between the working electrode and a reference electrode of the biosensor; and measuring the raw current generated thereby, wherein the electronics unit comprises a direct current measuring unit configured for measuring the raw current.
13. A system for operating a biosensor for electrochemically detecting an analyte value in a sample of a body fluid, the system comprising a biosensor operable by performing a method according to claim 1.
Description
BRIEF DESCRIPTION OF THE DRAWINGS
(1) The above-mentioned aspects of exemplary embodiments will become more apparent and will be better understood by reference to the following description of the embodiments taken in conjunction with the accompanying drawings, wherein:
(2) Further details of this disclosure may be derived from the following disclosure of preferred embodiments. The features of the embodiments may be realized in an isolated way or in any combination. This disclosure is not restricted to the embodiments. The embodiments are schematically depicted in the figures. Identical reference numbers in the figures refer to identical elements or functionally identical elements or elements corresponding to each other with regard to their functions.
(3) FIG. 1 schematically illustrates an electrical circuit being adapted for determining a sensitivity of a biosensor;
(4) FIGS. 2A-2B illustrate schematic mechanisms for measuring a sensitivity of a biosensor (FIG. 2A) and for a dielectric characterization of a biosensor (FIG. 2B), respectively;
(5) FIGS. 3A-3C illustrate an application of a potential step to the biosensor (FIG. 3A) and corresponding courses of a current response (FIG. 3B) and a related charge (FIG. 3C) of the biosensor;
(6) FIG. 4 illustrates a depiction of a corresponding course of the impedance of the biosensor in a Bode plot visualizing a frequency behavior of the biosensor;
(7) FIGS. 5A-5E illustrate a temporal course of a sensitivity (FIG. 5A), of an admittance (FIG. 5B), of a sensitivity-to-admittance ratio (FIG. 5C), of a relative deviation of the sensitivity-to-admittance ratio from a median (FIG. 5D), and of a capacitance (FIG. 5E) of the biosensor;
(8) FIGS. 6A-6C illustrate a temporal course of the current (FIG. 6A), the admittance (FIG. 6B), and of a current-to-admittance ratio (FIG. 6C) in a biosensor;
(9) FIG. 7 illustrates a schematic circuit diagram of the system comprising a biosensor and an electronics device;
(10) FIG. 8 illustrates a preferred example of a circuit especially adapted to charge determination; and
(11) FIGS. 9A-9C illustrate three preferred examples of circuits especially adapted to peak determination.
DESCRIPTION
(12) The embodiments described below are not intended to be exhaustive or to limit the invention to the precise forms disclosed in the following detailed description. Rather, the embodiments are chosen and described so that others skilled in the art may appreciate and understand the principles and practices of this disclosure.
(13) FIG. 1 schematically illustrates a number of aspects related to determining a sensitivity S of a biosensor 110. For a purpose of characterizing the biosensor 110 which constitutes an electrochemical cell as a whole, an electrical circuit 112 as schematically depicted in FIG. 1 may be applicable. Herein, a potentiostat 114 is employed, wherein the potentiostat 114 comprises outputs 116 which are each concurrently connected to one of the electrodes 118 of the biosensor 110, i.e., to a working electrode 120, a reference electrode 122, and a counter electrode 124. The potentiostat 114 may be adapted for adjusting and/or measuring an electrical potential difference between two of the electrodes 118 in the biosensor 110, in particular, between the working electrode 120 and the reference electrode 122. For this purpose, the potentiostat 114 may be implemented in order to be capable of injecting a current into the biosensor 110 through the counter electrode 124. The electrical circuit 112 may, thus, allow both adjusting the electrical potential difference between the working electrode 120 and the reference electrode 122 and, alternatively or in addition, measuring a direct raw current I between the working electrode 120 and the counter electrode 124. As a result, the electrical circuit 112 may be capable of measuring the raw current I between the working electrode 120 and the counter electrode 124.
(14) According to Equation (1),
S=(I−I.sub.0)/c, (1)
wherein the term I.sub.0 refers to a possible zero current, the sensitivity S of the biosensor 110 may, further, be obtained from a course of the direct raw current I with respect to a concentration c of an analyte, such as glucose, to be determined by the biosensor 110. Thus, the electrical circuit 112 may be capable of providing an overall response of the biosensor 110 to an analyte profile, such as a glucose profile, as applied to the biosensor 110. However, the DC raw current I cannot differentiate between effects which may arise from different partitions of the biosensor 110 as described below in more detail. In the electrical circuit 112, additional electrochemical techniques for detecting artefacts can only be applied to the working electrode 120 while artefacts related to the reference electrode 122 or the counter electrode 124 may remain undetectable hereby.
(15) FIG. 2A illustrates, in a highly schematic manner, a particularly preferred mechanism of an in-vivo determination of the sensitivity S of the biosensor 110, which may also be referred to as a “functional testing” of the biosensor 110. In the biosensor 110, the working electrode 120 having a surface area A may, typically, be placed on a substrate 126, preferably on a flexible printed circuit board 128, and be furnished with solder resists 130. Further, the working electrode 120 is covered by a membrane 132 having a thickness d. Herein, the membrane 132 may, preferably, comprise an enzyme 134, in particular glucose oxidase, often abbreviated to “GOD.” A reaction of an analyte 136, in particular glucose, and oxygen 138 as provided by the body fluid 140 may lead to a formation of hydrogen peroxide H.sub.2O.sub.2 which may react with manganese dioxide MnO.sub.2 also being present at the surface of the working electrode 120 as catalyst and/or mediator, thereby providing free electrons 2 e.sup.− to the working electrode 120, whereby the direct raw current I is generated. According to Equation (3),
S=P.sub.ana/d.Math.A, (3)
apart from the surface area A of the working electrode 120 and the thickness d of the membrane, a permeability P.sub.ana of the membrane 132 with respect to the analyte, such as glucose, may be capable of influencing the sensitivity S of the biosensor 110. As result, the functional testing of the biosensor 110 may provide the sensitivity S of the biosensor 110 which may depend from a number of variables, such as the thickness d and the area of the membrane 132 which may be varying due to manufacturing effects.
(16) FIG. 2B illustrates in a highly schematic manner a particularly preferred mechanism of a measurement of an in-vivo current response indicative of an in-vivo admittance Y(t) of the biosensor 110, which may also be referred to as an in-vivo “dielectric characterization” or a “detection of in-vivo properties” of the biosensor 110. Again, the working electrode 120 of the biosensor 110 having the surface area A may, typically, be placed on the substrate 126, such as the flexible printed circuit board 128, and be furnished with solder resists 130. As particularly preferred, the working electrode 120 may be covered by the membrane 132 having a thickness d. Again, the membrane 132 may, preferably, comprise the enzyme 134, in particular glucose oxidase. According to Equation (4),
Y(t)=˜P.sub.ion/d.Math.A, (4)
the admittance Y(t) of the biosensor 110 may depend on a permeability P.sub.ion of the membrane with respect to the ions, such as Na.sup.+ or Cl.sup.− ions, the thickness d of the membrane, and the area A of the electrode 118.
(17) As further indicated in FIG. 2B, the surface area A of the electrode 118 may be described by having a double layer being represented by a double-layer capacitance as schematically depicted in FIG. 8 below, wherein the double-layer capacitance may be determined by measuring the in-vivo current response of the biosensor 110. As used herein, the double-layer capacitance may be used as a quantity representing the surface area A of the electrode 118. A measurement of the double-layer capacitance may reveal changes related to electrode surface, in particular, loss of contact, draining, or detaching of the electrode 118. As a result, the measurement of the double-layer capacitance may be employed as additional parameter, particularly, adapted to provide additional failsafe information with regard to the operation of the biosensor.
(18) By comparing the respective results as schematically illustrated in FIGS. 2A and 2B, a sensitivity-to-admittance ratio S(t)/Y(t) may be determined which, advantageously, only depends on a ratio of the respective membrane permeabilities P.sub.ana, P.sub.ion with respect to the analyte and the ions in accordance with Equation (5):
S(t)/Y(t)=˜P.sub.ana/P.sub.ion (5)
(19) As described above, the determined sensitivity-to-admittance ratio S(t)/Y(t) may allow providing information about a current state of the intrinsic membrane transport properties related to the respective permeabilities of the membrane 132 while the geometric properties of the biosensor, in particular the thickness d of the membrane 132 and the geometric area A of the working electrode 120, can be disregarded. As a result, by determining the sensitivity-to-admittance ratio S(t)/Y(t), a change of the thickness d of the membrane 132, such as by a swelling of the membrane 132 during an in-vivo operation of the biosensor 110, can be disregarded.
(20) FIG. 3 illustrates an application of a potential step 150 to the biosensor 110 and a response of the biosensor 110 to the application of the potential step 150 as a preferred embodiment configured for determining the in-vivo current response indicative of the in-vivo admittance Y(t) of the biosensor.
(21) As schematically depicted in FIG. 3A, the potential step 150 can be considered as the application of an enhanced electrical potential E.sub.2 over a time interval Δt=t.sub.1−t.sub.0 with respect to the electrical potential E.sub.1 prevailing at the membrane, thus providing an electrical potential difference ΔE to the membrane over the time interval Δt. As an alternative (not depicted here), a diminished electrical potential E.sub.2 may be applied over the time interval Δt with respect to the electrical potential E.sub.1 prevailing at the membrane, again, thus providing an electrical potential difference ΔE to the membrane over the time interval Δt. Further alternatives may use a different time-varying electrical potential, in particular, a time-varying waveform, at least one linear or non-linear sweep, or at least one cyclically varying signal, such as described above in more detail. For sake of simplicity, the potential step 150 will include any of these time-varying electrical potentials in the following.
(22) FIG. 3B schematically shows a corresponding course 152 of a current response I(t) of the biosensor 110 as affected by a first application of a first potential step to the biosensor 110 at the time t.sub.0=0 s and, subsequently, a second application of a second potential step to the biosensor 110 at the time t.sub.1=0.24 s, whereby, in this particular example, the second application exhibits a reversed sign of the second potential step with respect to the first application of the first potential step. However, other kinds applications of potential steps are feasible, apart from varying the sign of the potential step 150 the height of the electrical potential difference ΔE may, alternatively or in addition, also be varied.
(23) Taking hereby a capacitance C of the membrane 132 into consideration, the current I(t) at the membrane 132 after the application of the potential step 150 may, as schematically depicted in FIG. 3B, follow exhibit an exponential decay 154 which can, after the first application of the first potential step which exhibits a positive sign, be described by any one of Equations (6) or (7):
(24)
wherein I.sub.max denotes a maximum current and I.sub.0 the zero current. For a negative sign of the potential step 150, the current I(t) at the membrane 132 after the second application of the second potential step can similarly be described with alternating signs.
(25) As further indicated in FIG. 3B and Equation (8), the exponential decay 154 can be described by referring to a term
τ=R.sub.M.Math.C, (8)
wherein the term τ relates to a time constant τ which may be assigned to the exponential decay 154 of the current I(t) in consequence of the application of the potential step 150 to the biosensor 110. As generally used, the time constant τ may be defined as relating to a time interval after which an initial intensity at the beginning of the time interval has decreased to a value of approximately l/e≈0.367879 of the initial intensity. However, other kinds of definitions for the time constant τ may also be applicable, such as a decay of the intensity after the time interval to a value of approximately ½ of the initial intensity.
(26) In particular, the exponential decay 154 as schematically depicted in FIG. 3B may, thus, be used for determining the electrical resistance R.sub.M of the membrane 132 according to Equation (9)
R.sub.M=ΔE/I.sub.max, (9)
whereby only the height of the electrical potential difference ΔE as applied to the biosensor 110 during the potential step 150 and the observed maximum current I.sub.max which can be derived from the course 152 of a current response I(t) of the biosensor 110 at the first operating point 156 below the time constant τ, preferably at a time interval of 10 μs to 100 μs after the application of the potential step 150.
(27) As further schematically depicted in FIG. 3B, a second operating point 158 is, in addition, selected above τ, preferably above 2τ, 3τ, 4τ, or 5τ, for determining the electrical capacity C of the working electrode 120. By applying the general definition of the capacitance C according to Equation (11)
C=![custom character]()
/ΔE, (11)
this may allow determining the additional charge
![custom character]()
(t)=∫I(t)dt, (12)
which has been provided to the membrane 132 by application of the potential step 150.
(28) FIG. 3C schematically depicts a corresponding course 160 of the additional charge ![custom character]()
(t) of the biosensor 110 as affected by the first application of the first potential step to the biosensor 110 at the time t.sub.0=0 s and, subsequently, the second application of the second potential step to the biosensor 110 at the time t.sub.1=0.24 s, whereby, in this particular example, the second application, again, exhibits the reversed sign of the second potential step with respect to the first application of the first potential step.
(29) FIG. 4 schematically depicts a “Bode plot” which, usually, describes a combination of a Bode magnitude plot referring to intensity versus an applied frequency ƒ and a Bode phase plot referring to a phase shift versus the applied frequency ƒ. As shown in FIG. 4 on the left-hand side, a logarithm of the absolute value of the impedance Z in Ohm and, on the right-hand side, a phase shift of the response of the biosensor 110 is plotted versus the logarithm of the frequency ƒ with respect to the base 10 of the alternating electric voltage or current as applied to biosensor 110. In FIG. 4, various curves 162 refer to the Bode magnitude plot related to the logarithm of the absolute value of the impedance Z versus the logarithm of the frequency ƒ.
(30) As can be further seen in FIG. 4, the curves 162 exhibit various features which may occur at predefined frequency ranges. On one hand, an increase 164 of the impedance Z observable towards lower frequencies is, usually, considered to be attributable to a capacitive behavior of the double layer C.sub.DL as described above with reference to FIG. 2B. On the other hand, a decrease 166 of the impedance Z observable towards higher frequencies is, usually, considered to be attributable to a high-frequency Ohmic behavior of the membrane resistance.
(31) As further disclosed by FIG. 4, the curves 162 exhibit a distinction 168 with respect to each other, particularly, in a range of 1 Hz to 10 kHz, in particular of 3 Hz to 3 kHz, especially of 10 Hz to 1 kHz. This behavior which expresses an alteration 170 of the electrical resistance of the membrane 132 can, generally, be attributed to an alteration of the permeability and thickness of the membrane 132, such as in consequence of a swelling of the membrane 132 during the in-vivo operation of the biosensor 110 as described above. Thus, it may be particularly be advantageous to measure the impedance Z of the biosensor 110 by application of a single frequency in the indicated range.
(32) FIG. 5 illustrates temporal courses of a number of quantities related to the biosensor 110 which may be provided by the measurements as described herein.
(33) Firstly, FIG. 5A illustrates the temporal course of the current response I(t) of the biosensor 110 at a constant concentration c of the analyte which is, according to Equation (1), proportional to the sensitivity S of the biosensor 110. As can be derived from FIG. 5A, a large sensitivity change accumulating up to 100% may occur, in particular, due to a swelling of the membrane 132 as, for example, expressed by Equation (3). As a result, the sensitivity S of the biosensor 110 is receptive to the operation of the biosensor 110 and, thus, not suited for determining an in-vivo drift in the biosensor 110 even when the concentration c of the analyte may stay constant.
(34) Similarly, FIG. 5B illustrates temporal courses of the admittance Y(t) of the biosensor 110, wherein curve 172 was obtained by application of a potential step 150 while curve 174 was obtained by application of electrochemical impedance spectroscopy (EIS), in particular for purposes of comparison. Irrespective of a manner of generation of the curves 172, 174, the admittance Y of the biosensor 110 depends on the geometric properties of the biosensor 110 since it changes its value due to the swelling of the membrane 132 as, for example, expressed by Equation (4).
(35) In contrast hereto, FIG. 5C illustrates temporal courses of the sensitivity-to-admittance ratio S(t)/Y(t) of the biosensor 110, which, according to Equation (5), do not depend on the geometric properties of the biosensor, in particular neither on the thickness d of the membrane 132 nor on the surface area A of the working electrode 120. Again, curve 172 was obtained by application of a potential step 150 while curve 174 was obtained by application of EIS. As a result, the sensitivity-to-admittance ratio S(t)/Y(t) of the biosensor 110 allows providing information about a current state of the intrinsic membrane transport properties related to the permeabilities P.sub.ana, P.sub.ion of the membrane 132 with respect to the analyte and to the ions. As can be derived from FIG. 5C, as long as the intrinsic membrane transport properties stays constant, the temporal course of the sensitivity-to-admittance ratio S(t)/Y(t) of the biosensor 110 remains unaffected by other changes of the membrane 132, such as the swelling of the membrane 132 over the time interval as depicted here. Consequently, the sensitivity-to-admittance ratio S(t)/Y(t) of the biosensor 110 as depicted in FIG. 5C, thus, allows determining the in-vivo drift of the biosensor 110 which is, subsequently, compensated when determining the analyte value by using the raw current.
(36) As a kind of enlargement of FIG. 5C, FIG. 5D illustrates temporal courses of a relative deviation of the sensitivity-to-admittance S(t)/Y(t) ratio from a median given in percent of the deviation from the median, wherein, again, curve 172 was obtained by application of a potential step 150 while curve 174 was obtained by application of EIS. As can be derived from FIG. 5D, the relative deviation of the sensitivity-to-admittance S(t)/Y(t) ratio from the median remains constant over the depicted time interval apart from time periods 176 in which the temperature of the membrane 132 in the biosensor 110 slightly varies. In fact, the variations of the temperature can be considered particularly small since they are too small to attract attention in FIG. 5C. This kind of behavior, thus, clearly demonstrates that the determination of the sensitivity-to-admittance S(t)/Y(t) ratio appears a reasonable quantity particularly suited for determining the in-vivo drift of the biosensor 110 since a temperature change may be considered as a factor triggering an in-vivo drift of the biosensor 110.
(37) As an alternative measure, FIG. 5E illustrates temporal courses of the capacitance C of the biosensor 110, again, showing curve 172 which was obtained by application of a potential step 150 while curve 174 was obtained by application of EIS. Similar to FIG. 5C, the temporal course of the capacitance C of the biosensor 110 stays practically constant over the depicted time interval.
(38) FIG. 6 presents a further example of temporal courses of a number of in-vivo properties related to the biosensor 110 which may be provided by the application of a potential step 150 s described herein, wherein, in contrast to FIGS. 5A to 5E, the time scale extends here over more than two and a half complete days.
(39) Herein, FIG. 6A illustrates the temporal course of the current response I(t) of the biosensor 110 at constant concentration c=10 mM of the analyte glucose. The corresponding admittance Y(t) of the biosensor 110 is depicted in FIG. 6B while a corresponding current-to-admittance ratio I(t)/Y(t) as shown in FIG. 6C is proportional to the sensitivity-to-admittance S(t)/Y(t) ratio of the biosensor 110 at a constant concentration of the analyte as applicable here. Again, from FIG. 6C it may be derived that, apart from the first hours of operation, the sensitivity-to-admittance S(t)/Y(t) ratio of the biosensor 110 remains constant within thresholds of ±5%, thus, implying here a perfectly compensated sensitivity drift of the biosensor 110.
(40) FIG. 7 illustrates a schematic circuit diagram of the system 200 comprising the biosensor 110 and an electronics unit 202, wherein the electronics unit 202 comprises a direct current measuring unit 204 and a potential step response measuring unit 206. Compared to other possible embodiments, the circuit of FIG. 7 comprises more analogue electronic elements which allow reducing the load on microcontrollers, thus, providing a faster processing within the electronics unit 202 with reduced technical effort.
(41) As depicted in FIG. 7, the direct current measuring unit 204 comprises an analog controller 208, which may control the potentiostat 114 as described above, which may be driven by an input 210, and which drives the electrodes 118, in particular the working electrode 120, the reference electrode 122, and the counter electrode 124, in particular by application of an electrical potential in order to measure the raw current I and, in addition, of the potential step 150 for measuring the in-vivo admittance the biosensor 110. Further, the direct current measuring unit 204 comprises a glucose current measuring unit 212, which is adapted to measure and to provide a DC output 214, which is the raw current I or a value related to the raw current I, preferably, a voltage converted raw current I, as measured for the analyte glucose. However, other kinds of values may also be provided at the DC output 214.
(42) As further shown in the exemplary embodiment of FIG. 7, the electronics unit 202 further comprises a number of switches 216 (four switches 216 are actually depicted here) which are configured to allow switching an output of the biosensor 110, in particular of the working electrode 120, between the glucose current measuring unit 212 as comprised by the direct current measuring unit 204 and one or more units as comprised by the potential step response measuring unit 206, in particular, to allow measuring the admittance of the biosensor 110 in addition to the raw current I.
(43) For this purpose, the potential step response measuring unit 206 may comprise a charge counter 218 which may provide a value related to the charge C accumulated in the membrane 132 of the working electrode 120 to a charge output 220. A preferred example of a circuit configured to be used as the charge counter 218 is shown in FIG. 8.
(44) Further, the potential step response measuring unit 206 may comprise a peak detector 222 which may provide information related to a peak the charge accumulated in the membrane 132 of the working electrode 120 to a peak information output 224, wherein the peak information may, preferably, be the maximum current I.sub.max or a value related hereto, in particular, a voltage converted maximum current I.sub.max. Three different exemplary embodiments of a circuit configured to be used as the peak detector 222 are shown in FIGS. 9A to 9C.
(45) According to the exemplary embodiment as depicted in FIG. 7, the potential step response measuring unit 206 may, in addition, comprise a fast sampling block 226, which may be configured to allow a fast sampling of the course 152 of the current response I(t) to the application of the potential step 150 to the biosensor 110. Herein, the course 152 of the current response I(t) may, thus, provide additional information that can be used in addition to the charge C and the maximum current I.sub.max as provided by the other two units 218, 222 of the potential step response measuring unit 206. In addition hereto, the potential step response measuring unit 206 may comprise further units for processing outputs as provided by the biosensor 110 and, hereby, acquiring additional information or the same information, in particular, for a purpose of redundancy.
(46) As mentioned above, FIG. 8 shows a preferred example of a circuit 228 for charge determination. As illustrated there, the circuit 228 comprises three successive stages 230, 232, 234, wherein each stage 230, 232, 234 has an operational amplifier. Herein, the first stage 230 is a current-voltage converter which provides the voltage-conversed course 152 of the current response I(t) at a connection point 236 after a resistor R24 as output. The second stage 232 is a differential amplifier while the third stage 234 is an integration unit which is configured to provide the desired value for the charge C at the output of the circuit 228.
(47) FIGS. 9A to 9C each illustrate a preferred example of a circuit 238 which is especially adapted to peak determination.
(48) As shown in FIG. 9A, the circuit 238 comprises three successive stages 240, 242, 244, wherein each stage 240, 242, 244 has an operational amplifier. Herein, the first stage 240 is, again, a current-voltage converter while the third stage 244 is, again, a differential amplifier. The second stage 242 comprises a combination of a capacitor C1 and a reverse-biased diode D1 which provide that incoming charges are stored in the capacitor C1 which cannot immediately be discharged due to the reverse-biased diode D1. As a result, a peak value can be determined by the combination of the capacitor C1 and the reverse-biased diode D1 which is, subsequently, be amplified in the third stage 244. An eventual discharge of the capacitor C1 can, according to this particular embodiment, only be achieved after a period of time. Only after this period of time a further peak value may be determined by using this particular embodiment of the circuit 238.
(49) Thus, in order to allow a faster repetition of measurements, the amended circuits 238 for peak determination as shown in FIGS. 9B and 9C may, preferably be used. Herein, the circuit 238 of FIG. 9B comprises a second stage 246 which has a combination of a diode D2, a capacitor C2 and a switch SW1, wherein the switch SW1 may be used for discharging the capacitor D2 if required. Further, the circuit 238 of FIG. 9C comprises an arrangement having four stages 240, 250, 248, 244 which allows an improved peak determination.
(50) As mentioned above, the present method further comprises monitoring a failsafe operation of the biosensor 110. For this purpose, a combination of at least two, preferably three measured values may be used. In particular, the following values may be considered to be related to corresponding technical parts and effects: the sensitivity S of the biosensor 110 may be related to an activity of the enzyme in the membrane 132, to an amount of catalyst and/or mediator in the membrane 132, and to a calibration value, particularly acquired by factory calibration or by initial calibration; the electrical resistance R.sub.M of the membrane 132 may, on one hand, be related to a swelling of the membrane 132 in-vivo (leading to a slow reaction when swelling) and, on the other hand, to a contact of the membrane 132 with the electrode material (leading to a fast reaction in case of loss); and the electrical capacitance C of the working electrode 120 may, on one hand, be related to an amount of catalyst and/or mediator at the working electrode 120 (leading to a slow reaction in case of loss) and, on the other hand, to a loss of contact of the working electrode 120 with the electrode pad (leading to a fast reaction in case of loss).
(51) Consequently, information concerning a behavior the sensitivity S of the biosensor 110 may be insufficient since they may be due to a number of different alterations within the biosensor 110. However, by combining the information concerning the behavior of the sensitivity S of the biosensor 110 with further information about the electrical resistance R.sub.M of the membrane 132 and the electrical capacitance C of the working electrode 120 may, nevertheless be capable for monitoring the failsafe operation of the biosensor 110, in particular, in accordance with the following Table. Herein, an availability of the sensitivity S of the biosensor 110 by in-vivo calibration may determine whether the information about the electrical resistance R.sub.M of the membrane 132 and the electrical capacitance C of the working electrode 120 can be used for compensation or as in a failsafe operation.
(52) As indicated in the Table, a possible reaction with regard to observations of change in at least one of the sensitivity S of the biosensor 110, the electrical resistance R.sub.M of the membrane 132, and the electrical capacitance C of the working electrode 120 may be selected from at least one of: an automatic “sensitivity-drift compensation;” an indication of a “no valid value;” a recommendation for “recalibration;” or a request for “shut-off” of the biosensor 110.
(53) While exemplary embodiments have been disclosed hereinabove, the present invention is not limited to the disclosed embodiments. Instead, this application is intended to cover any variations, uses, or adaptations of this disclosure using its general principles. Further, this application is intended to cover such departures from the present disclosure as come within known or customary practice in the art to which this invention pertains and which fall within the limits of the appended claims.
(54) TABLE-US-00001 Cause Occurrence S R C Possible reaction positive sensitivity beginning to mid of + − ∘ none OR drift due to wear-time (due to increased (due to increased (no change) sensitivity-drift membrane permeability of permeability of compensation swelling membrane for membrane for ions) glucose) negative sensitivity mid to end of wear − + o sensitivity-drift drift due to time (due to reduced mass (due to a reduced (no change) compensation OR encapsulation of transport to sensor current path for ion no valid value OR membrane due to encapsulation) conduction due to recalibration encapsulation) sudden membrane any time during wear ++ −− o no valid value OR defect time (due to far better (due to far better ion (no change) recalibration OR mass transport of conductivity as a shut-off glucose to enzyme) result of defect) sudden loss of any time during wear −− ∘/+ −− no valid value OR contact between time (due to less active (depending on (due to sudden recalibration OR paste electrode and/or less contacted electrode setup, decrease in area shut-off and pad area) constant OR sudden and/or mediator increase) and/or catalyst) slow loss of any time during wear − ∘ −− no valid value OR catalyst and/or time (after previous (constant) (due to slow loss of recalibration OR mediator reduction in C; due mediator and/or shut-off to less catalyst and/ catalyst) or mediator) slow loss of mid to end of wear − ∘ ∘ no valid value OR enzyme activity time (slow loss due to (constant) (constant) recalibration OR decrease in enzyme shut-off activity)
LIST OF REFERENCE NUMBERS
(55) 110 biosensor
(56) 112 electrical circuit
(57) 114 potentiostat
(58) 116 output
(59) 118 electrode
(60) 120 working electrode
(61) 122 reference electrode
(62) 124 counter electrode
(63) 126 substrate
(64) 128 printed circuit board
(65) 130 solder resist
(66) 132 membrane
(67) 134 enzyme
(68) 136 analyte
(69) 138 oxygen
(70) 140 body fluid
(71) 150 potential step
(72) 152 course of current response I(t)
(73) 154 exponential decay
(74) 156 first operating point
(75) 158 second operating point
(76) 160 course of additional charge ![custom character]()
(t)
(77) 162 curves in a Bode phase plot
(78) 164 increase towards lower frequencies
(79) 166 decrease towards higher frequencies
(80) 168 distinction
(81) 170 alteration of the electrical resistance
(82) 172 curve obtained by application of potential step
(83) 174 curve obtained by application of alternating current
(84) 200 system
(85) 202 electronics unit
(86) 204 direct current measuring unit
(87) 206 potential step response measuring unit
(88) 208 analog controller
(89) 210 input
(90) 212 glucose current measuring unit
(91) 214 DC output
(92) 216 switches
(93) 218 charge counter
(94) 220 charge output
(95) 222 peak detector
(96) 224 peak information output
(97) 226 fast sampling block
(98) 228 circuit for charge determination
(99) 230 stage
(100) 232 stage
(101) 234 stage
(102) 236 connection point
(103) 238 circuit for peak determination
(104) 240 stage
(105) 242 stage
(106) 244 stage
(107) 246 stage
(108) 248 stage
(109) 250 stage
