Patent Description:
Electrochemical glucose test strips, such as those used in the OneTouch® Ultra® whole blood testing kit, which is available from LifeScan, Inc. , are designed to measure the concentration of glucose in a physiological fluid sample from patients with diabetes. The measurement of glucose can be based on the selective oxidation of glucose by the enzyme glucose oxidase (GO). The reactions that can occur in a glucose test strip are summarized below in Equations <NUM> and <NUM>.

Eq. <NUM>     Glucose + GO(ox) → Gluconic Acid + GO(red).

Eq. <NUM>     GO(red) + <NUM> Fe(CN)<NUM><NUM>- → GO(ox) + <NUM> Fe(CN)<NUM><NUM>-.

As illustrated in Equation <NUM>, glucose is oxidized to gluconic acid by the oxidized form of glucose oxidase (GO(ox)). It should be noted that GO(ox) may also be referred to as an "oxidized enzyme. " During the reaction in Equation <NUM>, the oxidized enzyme GO(ox) is converted to its reduced state, which is denoted as GO(red) (i.e., "reduced enzyme"). Next, the reduced enzyme GO(red) is re-oxidized back to GO(ox) by reaction with Fe(CN)<NUM><NUM>- (referred to as either the oxidized mediator or ferricyanide) as illustrated in Equation <NUM>. During the re-generation of GO(red) back to its oxidized state GO(ox), Fe(CN)<NUM><NUM>- is reduced to Fe(CN)<NUM><NUM>- (referred to as either reduced mediator or ferrocyanide).

When the reactions set forth above are conducted with a test signal applied between two electrodes, a test current can be created by the electrochemical re-oxidation of the reduced mediator at the electrode surface. Thus, since, in an ideal environment, the amount of ferrocyanide created during the chemical reaction described above is directly proportional to the amount of glucose in the sample positioned between the electrodes, the test current generated would be proportional to the glucose content of the sample. A mediator, such as ferricyanide, is a compound that accepts electrons from an enzyme such as glucose oxidase and then donates the electrons to an electrode. As the concentration of glucose in the sample increases, the amount of reduced mediator formed also increases; hence, there is a direct relationship between the test current, resulting from the re-oxidation of reduced mediator, and glucose concentration. In particular, the transfer of electrons across the electrical interface results in the flow of a test current (<NUM> moles of electrons for every mole of glucose that is oxidized). The test current resulting from the introduction of glucose can, therefore, be referred to as a glucose signal.

Electrochemical biosensors may be adversely affected by the presence of certain blood components that may undesirably affect the measurement and lead to inaccuracies in the detected signal. This inaccuracy may result in an inaccurate glucose reading, leaving the patient unaware of a potentially dangerous blood sugar level, for example. As one example, the blood hematocrit level (i.e. the percentage of the amount of blood that is occupied by red blood cells) can erroneously affect a resulting analyte concentration measurement.

Variations in a volume of red blood cells within blood can cause variations in glucose readings measured with disposable electrochemical test strips. Typically, a negative bias (i.e., lower calculated analyte concentration) is observed at high hematocrit, while a positive bias (i.e., higher calculated analyte concentration as compared to referential analyte concentration) is observed at low hematocrit. At high hematocrit, for example, the red blood cells may impede the reaction of enzymes and electrochemical mediators, reduce the rate of chemistry dissolution since there is less plasma volume to solvate the chemical reactants, and slow diffusion of the mediator. These factors can result in a lower than expected glucose reading as less signal is produced during the electrochemical process. Conversely, at low hematocrit, fewer red blood cells may affect the electrochemical reaction than expected, and a higher measured signal can result. In addition, the physiological fluid sample resistance is also hematocrit dependent, which can affect voltage and/or current measurements.

Several strategies have been used to reduce or avoid hematocrit based variations on blood glucose. For example, test strips have been designed to incorporate meshes to remove red blood cells from the samples, or have included various compounds or formulations designed to increase the viscosity of red blood cells and attenuate the effect of low hematocrit on concentration determinations. Other test strips have included lysis agents and systems configured to determine hemoglobin concentration in an attempt to correct hematocrit. Further, biosensors have been configured to measure hematocrit by measuring an electrical response of the fluid sample via alternating current signals or change in optical variations after irradiating the physiological fluid sample with light, or measuring hematocrit based on a function of sample chamber fill time. These sensors have certain disadvantages. A common technique of the strategies involving detection of hematocrit is to use the measured hematocrit value to correct or change the measured analyte concentration, which technique is generally shown and described in the following respective <CIT>; <CIT>; <CIT>; <CIT>; <CIT>; <CIT>; <CIT>; <CIT>; <CIT>; <CIT>; <CIT>; or <CIT> and <CIT>.

For further background, <CIT> describes various embodiments that allow for determination of hematocrit by a time differential between the input and output signals such that a glucose measurement for a blood sample can be corrected by the measured hematocrit of a blood sample. <CIT> describes systems and exemplary methods of operating an analyte measurement system having a meter and a test strip. The methods and systems allow for trapping various errors during calculation of the analyte due to variations in the structure and materials making up the test strip and ambient temperatures. <CIT> describes a measurement with a test strip having two working electrodes, using the current transient for each working electrode measured at a predetermined durational offset from a peak of the current transient.

Applicant has devised systems and methods that allow for determination of an error in a sample fill condition for a biosensor. There is provided, according to the present invention, an analyte measurement system according to independent claim <NUM>. There is also provided an analyte measurement system according to independent claim <NUM>. There is further provided methods of determining a sample fill error in a biosensor according to independent claims <NUM> and <NUM>. Further optional features of the invention are recited in the dependent claims.

Accordingly, in any of the embodiments described earlier, the following features may also be utilized in various combinations with the previously disclosed embodiments. For example, the plurality of electrodes may include four electrodes with the first and second electrodes to measure the analyte concentration and third and fourth electrodes to measure the physical characteristic; the first, second, third and fourth electrodes are disposed in the same chamber provided on the substrate; the first and second electrodes and third and fourth electrodes are disposed in respective two different chambers provided on the substrate; all of the electrodes are disposed on the same plane defined by the substrate; a reagent is disposed proximate the at least two other electrodes and no reagent is disposed on the at least two electrodes; the final analyte concentration is determined from the second signal within about <NUM> seconds of a start of the test sequence and the predetermined threshold may include any value from about <NUM> to about <NUM>; the sampling time point is selected from a look-up table that includes a matrix in which different qualitative categories of the estimated analyte are set forth in the leftmost column of the matrix and different qualitative categories of the measured or estimated physical characteristic are set forth in the topmost row of the matrix and the sampling times are provided in the remaining cells of the matrix; the microcontroller determines the analyte concentration with an equation of the form: <MAT> where.

Moreover, in any of the embodiments described earlier, the following features may also be utilized in various combinations with the previously disclosed embodiments. For example, the microcontroller estimates the analyte concentration with an equation of the form: <MAT> where.

in which the microcontroller determines the analyte concentration with an equation of the form: <MAT> where:.

These and other embodiments, features and advantages will become apparent to those skilled in the art when taken with reference to the following more detailed description of the exemplary embodiments of the invention in conjunction with the accompanying drawings that are first briefly described.

The accompanying drawings, which are incorporated herein and constitute part of this specification, illustrate presently preferred embodiments of the invention, and, together with the general description given above and the detailed description given below, serve to explain features of the invention (wherein like numerals represent like elements), in which:.

As used herein, "oscillating signal" includes voltage signal(s) or current signal(s) that, respectively, change polarity or alternate direction of current or are multi-directional. Also used herein, the phrase "electrical signal" or "signal" is intended to include direct current signal, alternating signal or any signal within the electromagnetic spectrum. The terms "processor"; "microprocessor"; or "microcontroller" are intended to have the same meaning and are intended to be used interchangeably. As used herein, the term "annunciated" and variations on its root term indicate that an announcement may be provided via text, audio, visual or a combination of all modes or mediums of communication to a user.

<FIG> illustrates a test meter <NUM> for testing analyte (e.g., glucose) levels in the blood of an individual with a biosensor produced by the methods and techniques illustrated and described herein. Test meter <NUM> may include user interface inputs (<NUM>, <NUM>, <NUM>), which can be in the form of buttons, for entry of data, navigation of menus, and execution of commands. Data can include values representative of analyte concentration, and/or information that are related to the everyday lifestyle of an individual. Information, which is related to the everyday lifestyle, can include food intake, medication use, the occurrence of health check-ups, general health condition and exercise levels of an individual. Test meter <NUM> can also include a display <NUM> that can be used to report measured glucose levels, and to facilitate entry of lifestyle related information.

Test meter <NUM> may include a first user interface input <NUM>, a second user interface input <NUM>, and a third user interface input <NUM>. User interface inputs <NUM>, <NUM>, and <NUM> facilitate entry and analysis of data stored in the testing device, enabling a user to navigate through the user interface displayed on display <NUM>. User interface inputs <NUM>, <NUM>, and <NUM> include a first marking <NUM>, a second marking <NUM>, and a third marking <NUM>, which help in correlating user interface inputs to characters on display <NUM>.

Test meter <NUM> can be turned on by inserting a biosensor <NUM> (or its variants) into a strip port connector <NUM>, by pressing and briefly holding first user interface input <NUM>, or by the detection of data traffic across a data port <NUM>. Test meter <NUM> can be switched off by removing biosensor <NUM> (or its variants), pressing and briefly holding first user interface input <NUM>, navigating to and selecting a meter off option from a main menu screen, or by not pressing any buttons for a predetermined time. Display <NUM> can optionally include a backlight.

In one embodiment, test meter <NUM> of the analyte measurement system can be configured to not receive a calibration input for example, from any external source, when switching from a first test strip batch to a second test strip batch. Thus, in one exemplary embodiment, the meter is configured to not receive a calibration input from external sources, such as a user interface (such as inputs <NUM>, <NUM>, <NUM>), an inserted test strip, a separate code key or a code strip, data port <NUM>. Such a calibration input is not necessary when all of the biosensor batches have a substantially uniform calibration characteristic. The calibration input can be a set of values ascribed to a particular biosensor batch. For example, the calibration input can include a batch "slope" value and a batch "intercept" value for a particular biosensor batch. The calibrations input, such as batch slope and intercept values, may be preset within the meter as will be described below.

Referring to <FIG>, an exemplary internal layout of test meter <NUM> of the analyte measuring system is shown. Test meter <NUM> may include a processor <NUM>, which in some embodiments described and illustrated herein is a <NUM>-bit RISC microcontroller. In the preferred embodiments described and illustrated herein, processor <NUM> is preferably selected from the MSP <NUM> family of ultra-low power microcontrollers manufactured by Texas Instruments of Dallas, Texas. The processor can be bidirectionally connected via I/O ports <NUM> to a memory <NUM>, which in some embodiments described and illustrated herein is an EEPROM. Also connected to processor <NUM> via I/O ports <NUM> are the data port <NUM>, the user interface inputs <NUM>, <NUM>, and <NUM>, and a display driver <NUM>. Data port <NUM> can be connected to processor <NUM>, thereby enabling transfer of data between memory <NUM> and an external device, such as a personal computer. User interface inputs <NUM>, <NUM>, and <NUM> are directly connected to processor <NUM>. Processor <NUM> controls display <NUM> via display driver <NUM>. Memory <NUM> may be pre-loaded with calibration information, such as batch slope and batch intercept values, during production of test meter <NUM>. This pre-loaded calibration information can be accessed and used by processor <NUM> upon receiving a suitable signal (such as current) from the strip via strip port connector <NUM> so as to calculate a corresponding analyte level blood glucose concentration) using the signal and the calibration information without receiving calibration input from any external source.

Referring to <FIG> AND 2C THROUGH <NUM>, another embodiment of a hand-held test meter <NUM> of an analyte measurement system is provided. This version of the meter <NUM> includes a display <NUM>, a plurality of user interface buttons <NUM>, a strip port connector <NUM>, a USB interface <NUM>, and a housing. Referring to <FIG> AND 2C in particular, hand-held test meter <NUM> also includes a microcontroller block <NUM>, a physical characteristic measurement block <NUM>, a display control block <NUM>, a memory block <NUM> and other electronic components (not shown) for applying a test voltage to biosensor, and also for measuring an electrochemical response (e.g., plurality of test current values) and determining an analyte based on the electrochemical response. To simplify the current descriptions, the FIGUREs do not depict all such electronic circuitry.

Display <NUM> can be, for example, a liquid crystal display or a bi-stable display configured to show a screen image. An example of a screen image may include a glucose concentration, a date and time, an error message, and a user interface for instructing an end user how to perform a test.

Strip port connector <NUM> is configured to operatively interface with a biosensor <NUM>, such as an electrochemical-based biosensor configured for the determination of glucose in a whole blood sample. Therefore, the biosensor is configured for operative insertion into strip port connector <NUM> and to operatively interface with phase-shift-based hematocrit measurement block <NUM> via, for example, suitable electrical contacts.

USB Interface <NUM> can be any suitable interface known to one skilled in the art. USB Interface <NUM> is essentially a passive component that is configured to power and provide a data line to hand-held test meter <NUM>.

Once a biosensor is interfaced with hand-held test meter <NUM>, or prior thereto, a bodily fluid sample (e.g., a whole blood sample) is introduced into a sample chamber of the biosensor. The biosensor can include enzymatic reagents that selectively and quantitatively transform an analyte into another predetermined chemical form. For example, the biosensor can include an enzymatic reagent with ferricyanide and glucose oxidase so that glucose can be physically transformed into an oxidized form.

Memory block <NUM> of hand-held test meter <NUM> includes a suitable algorithm and can be configured, along with microcontroller block <NUM> to determine an analyte based on the electrochemical response of biosensor and the hematocrit of the introduced sample. For example, in the determination of the analyte blood glucose, the hematocrit can be used to compensate for the effect of hematocrit on electrochemically determined blood glucose concentrations.

Microcontroller block <NUM> is disposed within housing and can include any suitable microcontroller and/or micro-processer known to those of skill in the art. One such suitable microcontroller is a microcontroller commercially available from Texas Instruments, Dallas, TX USA and part number MSP430F5138. This microcontroller can generate a square wave of <NUM> to <NUM> and a <NUM> degree phase-shifted wave of the same frequency and, thereby, function as a signal generation s-block described further below. MSP430F5138 also has Analog-to-Digital (A/D) processing capabilities suitable for measuring voltages generated by phase shift based hematocrit measurement blocks employed in embodiments of the present disclosure.

Referring in particular to <FIG>, phase-shift-based hematocrit measurement block <NUM> includes a signal generation sub-block <NUM>, a low pass filter sub-block <NUM>, an biosensor sample cell interface sub-block <NUM>, an optional calibration load block <NUM> (within the dashed lines of <FIG>), a transimpedance amplifier sub-block <NUM>, and a phase detector sub-block <NUM>.

As described further below, phase-shift-based hematocrit measurement block <NUM> and microcontroller block <NUM> are configured to measure the phase shift of a bodily fluid sample in a sample cell of an biosensor inserted in the hand-held test meter by, for example, measuring the phase shift of one or more high frequency electrical signals driven through the bodily fluid sample. In addition, microcontroller block <NUM> is configured to compute the hematocrit of the bodily fluid based on the measured phase shift. Microcontroller <NUM> can compute the hematocrit by, for example, employing an A/D converter to measure voltages received from a phase-detector sub-block, convert the voltages into a phase-shift and then employing a suitable algorithm or look-up table to convert the phase-shit into a hematocrit value. Once apprised of the present disclosure, one skilled in the art will recognize that such an algorithm and/or look-up table will be configured to take into account various factors such as strip geometry (including electrode area and sample chamber volume) and signal frequency.

It has been determined that a relationship exists between the reactance of a whole blood sample and the hematocrit of that sample. Electrical modeling of a bodily fluid sample (i.e., a whole blood sample) as parallel capacitive and resistive components indicates that when an alternating current (AC) signal is forced through the bodily fluid sample, the phase shift of the AC signal will be dependent on both the frequency of the AC voltage and the hematocrit of the sample. Moreover, modeling indicates that hematocrit has a relatively minor effect on the phase shift when the frequency of the signal is in the range of approximately <NUM> to <NUM> and a maximum effect on the phase shift when the frequency of the signal is in the range of approximately <NUM> to <NUM>. Therefore, the hematocrit of a bodily fluid sample can be measured by, for example, driving AC signals of known frequency through the bodily fluid sample and detecting their phase shift. For example, the phase-shift of a signal with a frequency in the range of <NUM> to <NUM> can be used as a reference reading in such a hematocrit measurement while the phase shift of a signal with a frequency in the range of <NUM> to <NUM> can be used as the primary measurement.

Referring to <FIG>, in particular, signal generation sub-block <NUM> can be any suitable signal generation block and is configured to generate a square wave (0V to Vref) of a desired frequency. Such a signal generation sub-block can, if desired, be integrated into microcontroller block <NUM>.

The signal generated by signal generation sub-block <NUM> is communicated to dual low pass filter sub-block <NUM>, which is configured to convert the square wave signal to a sine wave signal of a predetermined frequency. The dual LPF of <FIG> is configured to provide both a signal of a first frequency (such as a frequency in the range of <NUM> to <NUM>) and a signal of a second frequency (such as a frequency in the range of <NUM> to <NUM>) to the biosensor sample cell interface sub-block and an biosensors' sample chamber (also referred to as the HCT measurement cell). Selection of the first and second frequency is accomplished using switch IC7 of <FIG>. The dual LPF of <FIG> includes employs two suitable operational amplifiers (IC4 and IC5) such as the operational amplifier available from Texas Instruments, Dallas, Texas, USA as high-speed, voltage feedback, CMOS operational amplifier part number OPA354.

Referring to <FIG>, F-DRV represents a square wave input of either a low or high frequency (e.g., <NUM> or <NUM>) and is connected to both IC4 and IC5. Signal Fi-HIGH/LOW (from the microcontroller) selects the output of dual low pass filter sub-block <NUM> via switch IC7. C5 in <FIG> is configured to block the operating voltage of dual low pass filter sub-block <NUM> from the HCT measurement cell.

Although a specific dual LPF is depicted in <FIG>, dual low pass filter sub-block <NUM> can be any suitable low pass filter sub-block known to one skilled in the art including, for example, any suitable multiple feedback low pass filter, or a Sallen and Key low pass filter.

The sine wave produced by low pass filter sub-block <NUM> is communicated to biosensor sample cell interface sub-block <NUM> where it is driven across the sample cell of the biosensor (also referred to as an HCT measurement cell). Biosensor sample cell interface block <NUM> can be any suitable sample cell interface block including, for example, an interface block configured to operatively interface with the sample cell of the biosensor via first electrode and second electrodes of the biosensor disposed in the sample cell. In such a configuration, the signal can be driven into the sample cell (from the low pass filter sub-block) via the first electrode and picked-up from the sample cell (by the transimpedance amplifier sub-block) via the second electrode as depicted in <FIG>.

The current produced by driving the signal across the sample cell is picked-up by transimpedance amplifier sub-block <NUM> and converted into a voltage signal for communication to phase detector sub-block <NUM>.

Transimpedance sub-block <NUM> can be any suitable transimpedance sub-block known to one skilled in the art. <FIG> is a simplified annotated schematic block diagram of one such transimpedance amplifier sub-block (based on two OPA354 operational amplifiers, IC3 and IC9). The first stage of TIA sub-block <NUM> operates at, for example, 400mV, which limits the AC amplitude to +/-400mV. The second stage of TIA sub-block <NUM> operates at Vref/<NUM>, a configuration which enables the generation of an output of the full span of the microcontroller A/D inputs. C9 of TIA sub-block <NUM> serves as a blocking component that only allows an AC sine wave signal to pass.

Phase detector sub-block <NUM> can be any suitable phase detector sub-block that produces either a digital frequency that can be read back by microcontroller block <NUM> using a capture function, or an analog voltage that can be read back by microcontroller block <NUM> using an analog to digital converter. <FIG> depicts a schematic that includes two such phase detector sub-blocks, namely an XOR phase detector (in the upper half of <FIG> and including IC22 and IC23) and a Quadrature DEMUX phase detector (in the lower half of <FIG> and including IC12 and IC13).

<FIG> also depicts a calibration load sub-block <NUM> that includes a switch (IC16) and a dummy load R7 and C6. Calibration load sub-block <NUM> is configured for the dynamic measurement of a phase offset for the known phase shift of zero degrees produced by resistor R7, thus providing a phase offset for use in calibration. C6 is configured to force a predetermined slight phase shift, e.g. to compensate for phase delays caused by parasitic capacities in the signal traces to the sample cell, or for phase delays in the electrical circuits (LPF and TIA).

The Quadrature DEMUX phase detector circuit of <FIG> includes two portions, one portion for a resistive part of the incoming AC signal and one portion for the reactive portion of the incoming AC signal. Use of such two portions enables the simultaneous measurement of both the resistive and reactive portion of the AC signal and a measurement range that covers <NUM> degrees to <NUM> degrees. The Quadrature DEMUX circuit of <FIG> generates two separate output voltages. One of these output voltages represents the "in phase measurement" and is proportional to the "resistive" part of the AC signal, the other output voltage represents the "Quadrature Measurement" and is proportional to the "reactive part of the signal. The phase shift is calculated as: <MAT>.

Such a Quadrature DEMUX phase detector circuit can also be employed to measure the impedance of a bodily fluid sample in the sample cell. It is hypothesized, without being bound, that the impedance could be employed along with the phase-shift, or independently thereof, to determine the hematocrit of the bodily sample. The amplitude of a signal forced through the sample cell can be calculated using the two voltage outputs of the Quadrature DEMUX circuit as follows : <MAT>.

This amplitude can then be compared to an amplitude measured for the known resistor of calibration load block <NUM> to determine the impedance.

The XOR phase detector portion has a measurement range of <NUM>° to <NUM>°, or alternatively a measurement range of -<NUM>° to +<NUM>°, depending whether the "Square wave input from µC" is in phase to the sine wave or is set to a <NUM>° phase shift. The XOR phase detector produces an output frequency that is always double the input frequency, however the duty cycle varies. If both inputs are perfectly in phase, the output is LOW, if both inputs are <NUM>° shifted the output is always HIGH. By integrating the output signal (e.g. via a simple RC element) a voltage can be generated that is directly proportional to the phase shift between both inputs.

As provided herein, one skilled in the art will recognize that phase detector sub-blocks employed in embodiments of the present disclosure can take any suitable form and include, for example, forms that employ rising edge capture techniques, dual edge capture techniques, XOR techniques and synchronous demodulation techniques.

Since low pass filter sub-block <NUM>, transimpedance amplifier sub-block <NUM> and phase detector sub-block <NUM> can introduce a residual phase shift into phase-shift-based hematocrit measurement block <NUM>, calibration load block <NUM> can be optionally included in the phase-shift-based hematocrit measurement block. Calibration load block <NUM> is configured to be essentially resistive in nature (for example a <NUM>-ohm load) and, therefore, induces no phase shift between excitation voltage and generated current. Calibration load block <NUM> is configured to be switched in across the circuit to give a "zero" calibration reading. Once calibrated, the hand-held test meter can measure the phase shift of a bodily fluid sample, subtract the "zero" reading to compute a corrected phase shift and subsequently compute the physical characteristic of the sample based on the corrected phase shift.

<FIG> is an exemplary exploded perspective view of the test strip <NUM> of the analyte measurement system, which may include seven layers disposed on a substrate <NUM>. The seven layers disposed on substrate <NUM> can be a first conductive layer <NUM> (which can also be referred to as electrode layer <NUM>), an insulation layer <NUM>, two overlapping reagent layers 22a and 22b, an adhesive layer <NUM> which includes adhesive portions <NUM>, <NUM>, and <NUM>, a hydrophilic layer <NUM>, and a top layer <NUM> which forms a cover <NUM> for the test strip <NUM>. Test strip <NUM> may be manufactured in a series of steps where the conductive layer <NUM>, insulation layer <NUM>, reagent layers <NUM>, and adhesive layer <NUM> are sequentially deposited on substrate <NUM> using, for example, a screen-printing process. Note that the electrodes <NUM>, <NUM>, and <NUM> are disposed for contact with the reagent layer 22a and 22b whereas the physical characteristic sensing electrodes 19a and 20a are spaced apart and not in contact with the reagent layer <NUM>. Hydrophilic layer <NUM> and top layer <NUM> can be disposed from a roll stock and laminated onto substrate <NUM> as either an integrated laminate or as separate layers. Test strip <NUM> has a distal portion <NUM> and a proximal portion <NUM> as shown in <FIG>.

Test strip <NUM> includes a sample-receiving chamber <NUM> through which a physiological fluid sample <NUM> may be drawn through or deposited (<FIG>). The physiological fluid sample discussed herein may be blood. Sample-receiving chamber <NUM> can include an inlet at a proximal end and an outlet at the side edges of test strip <NUM>, as illustrated in <FIG>. A fluid sample <NUM> can be applied to the inlet along axis L-L (<FIG>) to fill a sample-receiving chamber <NUM> so that glucose can be measured. The side edges of a first adhesive pad <NUM> and a second adhesive pad <NUM> located adjacent to reagent layer <NUM> each define a wall of sample-receiving chamber <NUM>, as illustrated in <FIG>. A bottom portion or "floor" of sample-receiving chamber <NUM> may include a portion of substrate <NUM>, conductive layer <NUM>, and insulation layer <NUM>, as illustrated in <FIG>. A top portion or "roof' of sample-receiving chamber <NUM> may include distal hydrophilic portion <NUM>, as illustrated in <FIG>. For test strip <NUM>, as illustrated in <FIG>, substrate <NUM> can be used as a foundation for helping support subsequently applied layers. Substrate <NUM> can be in the form of a polyester sheet such as a polyethylene tetraphthalate (PET) material (Hostaphan PET supplied by Mitsubishi). Substrate <NUM> can be in a roll format, nominally <NUM> microns thick by <NUM> millimeters wide and approximately <NUM> meters in length.

A conductive layer is required for forming electrodes that are used for the electrochemical measurement of glucose. First conductive layer <NUM> can be made from a carbon ink that is screen-printed onto substrate <NUM>. In a screen-printing process, carbon ink is loaded onto a screen and then transferred through the screen using a squeegee. The printed carbon ink can be dried using hot air at about <NUM>. The carbon ink can include VAGH resin, carbon black, graphite (KS15), and one or more solvents for the resin, carbon and graphite mixture. More particularly, the carbon ink may incorporate a ratio of carbon black: VAGH resin of about <NUM>:<NUM> and a ratio of graphite: carbon black of about <NUM>:<NUM> in the carbon ink.

For test strip <NUM>, as illustrated in <FIG>, first conductive layer <NUM> may include a reference electrode <NUM>, a first working electrode <NUM>, a second working electrode <NUM>, third and fourth physical characteristic sensing electrodes 19a and 19b, a first contact pad <NUM>, a second contact pad <NUM>, a reference contact pad <NUM>, a first working electrode track <NUM>, a second working electrode track <NUM>, a reference electrode track <NUM>, and a strip detection bar <NUM>. The physical characteristic sensing electrodes 19a and 20a are provided with respective electrode tracks 19b and 20b. The conductive layer may be formed from carbon ink. First contact pad <NUM>, second contact pad <NUM>, and reference contact pad <NUM> may be adapted to electrically connect to a test meter. First working electrode track <NUM> provides an electrically continuous pathway from first working electrode <NUM> to first contact pad <NUM>. Similarly, second working electrode track <NUM> provides an electrically continuous pathway from second working electrode <NUM> to second contact pad <NUM>. Similarly, reference electrode track <NUM> provides an electrically continuous pathway from reference electrode <NUM> to reference contact pad <NUM>. Strip detection bar <NUM> is electrically connected to reference contact pad <NUM>. Third and fourth electrode tracks 19b and 20b connect to the respective electrodes 19a and 20a. A test meter can detect that test strip <NUM> has been properly inserted by measuring a continuity between reference contact pad <NUM> and strip detection bar <NUM>, as illustrated in <FIG>.

Variations of the test strip <NUM> of an analyte measurement system (<FIG>, <FIG>, <FIG>, or <FIG>) are shown in <FIG>. Briefly, with regard to variations of test strip <NUM> (illustrated exemplarily in <FIG>), these test strips include an enzymatic reagent layer disposed on the working electrode, a patterned spacer layer disposed over the first patterned conductive layer and configured to define a sample chamber within the biosensor, and a second patterned conductive layer disposed above the first patterned conductive layer. The second patterned conductive layer includes a first phase-shift measurement electrode and a second phase-shift measurement electrode. Moreover, the first and second phase-shift measurement electrodes are disposed in the sample chamber and are configured to measure, along with the hand-held test meter, a phase shift of an electrical signal forced through a bodily fluid sample introduced into the sample chamber during use of the biosensor. Such phase-shift measurement electrodes are also referred to herein as bodily fluid phase-shift measurement electrodes. Biosensors of various embodiments described herein are believed to be advantageous in that, for example, the first and second phase-shift measurement electrodes are disposed above the working and reference electrodes, thus enabling a sample chamber of advantageously low volume. This is in contrast to a configuration wherein the first and second phase-shift measurement electrodes are disposed in a co-planar relationship with the working and reference electrodes thus requiring a larger bodily fluid sample volume and sample chamber to enable the bodily fluid sample to cover the first and second phase-shift measurement electrodes as well as the working and reference electrodes.

In the embodiment of <FIG> which is a variation of the test strip of <FIG>, an additional electrode 10a is provided as an extension of any of the plurality of electrodes 19a, 20a, <NUM>, <NUM>, and <NUM>. It must be noted that the built-in shielding or grounding electrode 10a is used to reduce or eliminate any capacitance coupling between the finger or body of the user and the characteristic measurement electrodes 19a and 20a. The grounding electrode 10a allows for any capacitance to be directed away from the sensing electrodes 19a and 20a. To do this, the grounding electrode 10a can be connected any one of the other five electrodes or to its own separate contact pad (and track) for connection to ground on the meter instead of one or more of contact pads <NUM>, <NUM>, <NUM> via respective tracks <NUM>, <NUM>, and <NUM>. In a preferred embodiment, the grounding electrode 10a is connected to one of the three electrodes that has reagent <NUM> disposed thereon. In a most preferred embodiment, the grounding electrode 10a is connected to electrode <NUM>. Being the grounding electrode, it is advantageous to connect the grounding electrode to the reference electrode (<NUM>) so not to contribute any additional current to the working electrode measurements which may come from background interfering compounds in the sample. Further by connecting the shield or grounding electrode 10a to electrode <NUM> this is believed to effectively increase the size of the counter electrode <NUM> which can become limiting especially at high signals. In the embodiment of <FIG>, the reagent are arranged so that they are not in contact with the measurement electrodes 19a and 20a. Alternatively, in the embodiment of <FIG>, the reagent <NUM> is arranged so that the reagent <NUM> contacts at least one of the sensing electrodes 19a and 20a.

In alternate version of test strip <NUM>, shown here in <FIG>, the top layer <NUM>, hydrophilic film layer <NUM> and spacer <NUM> have been combined together to form an integrated assembly for mounting to the substrate <NUM> with reagent layer <NUM>' disposed proximate insulation layer <NUM>'.

In the embodiment of <FIG>, the analyte measurement electrodes <NUM>, <NUM>, and <NUM> are disposed in generally the same configuration as in <FIG>, <FIG>, or <FIG>. The electrodes 19a and 20a to sense a hematocrit level, however, are disposed in a spaced apart configuration in which one electrode 19a is proximate an entrance 92a to the test chamber <NUM> and another electrode 20a is at the opposite end of the test chamber <NUM>. Electrodes <NUM>, <NUM>, and <NUM> are disposed to be in contact with a reagent layer <NUM>.

In <FIG>, <FIG>, <FIG> and <FIG>, the hematocrit sensing electrodes 19a and 20a are disposed adjacent each other and may be placed at the opposite end 92b of the entrance 92a to the test chamber <NUM> (<FIG> and <FIG>) or adjacent the entrance 92a (<FIG> and <FIG>). In all of these embodiments, the physical characteristic sensing electrodes are spaced apart from the reagent layer <NUM> so that these physical characteristic sensing electrodes are not impacted by the electrochemical reaction of the reagent in the presence of a fluid sample (e.g., blood or interstitial fluid) containing glucose.

In the various embodiments of the biosensor, there are two measurements that are made to a fluid sample deposited on the biosensor. One measurement is that of the concentration of the glucose in the fluid sample while the other is that of hematocrit in the same sample. The measurement of the hematocrit is used to modify or correct the glucose measurement so as to remove or reduce the effect of red blood cells on the glucose measurements. Both measurements (glucose and hematocrit) can be performed in sequence, simultaneously or overlapping in duration. For example, the glucose measurement can be performed first then the hematocrit; the hematocrit measurement first then the glucose measurement; both measurements at the same time; or a duration of one measurement may overlap a duration of the other measurement. Each measurement is discussed in detail as follow with respect to <FIG> and <FIG>.

<FIG> is an exemplary chart of a test signal applied to test strip <NUM> and its variations shown here in FIGUREs 3A-3T. Before a fluid sample is applied to test strip <NUM> (or its variants), test meter <NUM> is in a fluid detection mode in which a first test signal of about <NUM> millivolts is applied between second working electrode and reference electrode. A second test signal of about <NUM> millivolts is preferably applied simultaneously between first working electrode (e.g., electrode <NUM> of strip <NUM>) and reference electrode (e.g., electrode <NUM> of strip <NUM>). Alternatively, the second test signal may also be applied contemporaneously such that a time interval of the application of the first test signal overlaps with a time interval in the application of the second test voltage. The test meter may be in a fluid detection mode during fluid detection time interval TFD prior to the detection of physiological fluid at starting time at zero. In the fluid detection mode, test meter <NUM> determines when a fluid is applied to test strip <NUM> (or its variants) such that the fluid wets either the first working electrode <NUM> or second working electrode <NUM> (or both working electrodes) with respect to reference electrode <NUM>. Once test meter <NUM> recognizes that the physiological fluid has been applied because of, for example, a sufficient increase in the measured test current at either or both of first working electrode <NUM> and second working electrode <NUM>, test meter <NUM> assigns a zero second marker at zero time "<NUM>" and starts the test time interval TS. Test meter <NUM> may sample the current transient output at a suitable sampling rate, such as, for example, every <NUM> milliseconds to every <NUM> milliseconds. Upon the completion of the test time interval TS, the test signal is removed. For simplicity, <FIG> only shows the first test signal applied to test strip <NUM> (or its variants).

Hereafter, a description of how analyte (e.g., glucose) concentration is determined from the known signal transients (e.g., the measured electrical signal response in nanoamperes as a function of time) that are measured when the test voltages of <FIG> are applied to the test strip <NUM> (or its variants).

In <FIG>, the first and second test voltages applied to test strip <NUM> (or its variants described herein) are generally from about +<NUM> millivolts to about +<NUM> millivolts. In one embodiment in which the electrodes include carbon ink and the mediator includes ferricyanide, the test signal is about +<NUM> millivolts. Other mediator and electrode material combinations will require different test voltages, as is known to those skilled in the art. The duration of the test voltages is generally from about <NUM> to about <NUM> seconds after a reaction period and is typically about <NUM> seconds after a reaction period. Typically, test sequence time TS is measured relative to time t<NUM>. As the voltage <NUM> is maintained in <FIG> for the duration of TS, output signals are generated, shown here in <FIG> with the current transient <NUM> for the first working electrode <NUM> being generated starting at zero time and likewise the current transient <NUM> for the second working electrode <NUM> is also generated with respect to the zero time. It is noted that while the signal transients <NUM> and <NUM> have been placed on the same referential zero point for purposes of explaining the process, in physical term, there is a slight time differential between the two signals due to fluid flow in the chamber towards each of the working electrodes <NUM> and <NUM> along axis L-L. However, the current transients are sampled and configured in the microcontroller to have the same start time. In <FIG>, the current transients build up to a peak proximate peak time Tp at which time, the current slowly drops off until approximately one of <NUM> seconds or <NUM> seconds after zero time. At the point <NUM>, approximately at <NUM> seconds, the output signal for each of the working electrodes <NUM> and <NUM> may be measured and added together. Alternatively, the signal from only one of the working electrodes <NUM> and <NUM> can be doubled.

Referring back to <FIG>, the system drives a signal to measure or sample the output signals IE from at least one the working electrodes (<NUM> and <NUM>) at any one of a plurality of time points or positions T<NUM>, T<NUM>, T<NUM>,. As can be seen in <FIG>, the time position can be any time point or interval in the test sequence TS. For example, the time position at which the output signal is measured can be a single time point T<NUM> at <NUM> seconds or an interval <NUM> (e.g., interval~<NUM> milliseconds or more depending on the sampling rate of the system) overlapping the time point T<NUM> proximate <NUM> seconds.

From knowledge of the parameters of the biosensor (e.g., batch calibration code offset and batch slope) for the particular test strip <NUM> and its variations, the analyte (e.g., glucose) concentration can be calculated. Output transient <NUM> and <NUM> can be sampled to derive signals IE (by summation of each of the current IWE1 and IWE2 or doubling of one of IWE1 or IWE2) at various time positions during the test sequence. From knowledge of the batch calibration code offset and batch slope for the particular test strip <NUM>, the analyte (e.g., glucose) concentration can be calculated.

It is noted that "Intercept" and "Slope" are the values obtained by measuring calibration data from a batch of biosensors. Typically around <NUM> biosensors are selected at random from the lot or batch. Physiological fluid (e.g., blood) from donors is spiked to various analyte levels, typically six different glucose concentrations. Typically, blood from <NUM> different donors is spiked to each of the six levels. Eight biosensors (or strips in this embodiment) are given blood from identical donors and levels so that a total of <NUM> x <NUM> x <NUM> = <NUM> tests are conducted for that lot. These are benchmarked against actual analyte level ( blood glucose concentration) by measuring these using a standard laboratory analyzer such as Yellow Springs Instrument (YSI). A graph of measured glucose concentration is plotted against actual glucose concentration (or measured current versus YSI current) and a formula y = mx+c least squares fitted to the graph to give a value for batch slope m and batch intercept c for the remaining strips from the lot or batch. The applicants have also provided methods and systems in which the batch slope is derived during the determination of an analyte concentration. The "batch slope", or "Slope", may therefore be defined as the measured or derived gradient of the line of best fit for a graph of measured glucose concentration plotted against actual glucose concentration (or measured current versus YSI current). The "batch intercept", or "Intercept", may therefore be defined as the point at which the line of best fit for a graph of measured glucose concentration plotted against actual glucose concentration (or measured current versus YSI current) meets the y axis.

It is worthwhile here to note that the various components, systems and procedures described earlier allow for applicant to provide an analyte measurement system that heretofore was not available in the art. In particular, this system includes a biosensor that has a substrate and a plurality of electrodes connected to respective electrode connectors. The system further includes an analyte meter <NUM> that has a housing, a test strip port connector configured to connect to the respective electrode connectors of the test strip, and a microcontroller <NUM>, shown here in <FIG>. The microcontroller <NUM> is in electrical communication with the test strip port connector <NUM> to apply electrical signals or sense electrical signals from the plurality of electrodes.

Referring to <FIG>, details of a preferred implementation of meter <NUM> where the same numerals in <FIG> and <FIG> have a common description. In <FIG>, a strip port connector <NUM> is connected to the analogue interface <NUM> by five lines including an impedance sensing line EIC to receive signals from physical characteristic sensing electrode(s), alternating signal line AC driving signals to the physical characteristic sensing electrode(s), reference line for a reference electrode, and signal sensing lines from respective working electrode <NUM> and working electrode <NUM>. A strip detection line <NUM> can also be provided for the connector <NUM> to indicate insertion of a test strip. The analog interface <NUM> provides four inputs to the processor <NUM>: (<NUM>) real impedance Z'; (<NUM>) imaginary impedance Z"; (<NUM>) signal sampled or measured from working electrode <NUM> of the biosensor or I we1; (<NUM>) signal sampled or measured from working electrode <NUM> of the biosensor or I we2. There is one output from the processor <NUM> to the interface <NUM> to drive an oscillating signal AC of any value from <NUM> to about <NUM> or higher to the physical characteristic sensing electrodes. A phase differential P (in degrees) can be determined from the real impedance Z' and imaginary impedance Z" where: <MAT>
and magnitude M (in ohms and conventionally written as | Z | ) from line Z' and Z" of the interface <NUM> can be determined where <MAT>.

In this system, the microprocessor is configured to: (a) apply a first signal to the plurality of electrodes so that a batch slope defined by a physical characteristic of a fluid sample is derived and (b) apply a second signal to the plurality of electrodes so that an analyte concentration is determined based on the derived batch slope. For this system, the plurality of electrodes of the test strip or biosensor includes at least two electrodes to measure the physical characteristic and at least two other electrodes to measure the analyte concentration. For example, the at least two electrodes and the at least two other electrodes are disposed in the same chamber provided on the substrate. Alternatively, the at least two electrodes and the at least two other electrodes are disposed in respective two different chambers provided on the substrate. It is noted that for some embodiments, all of the electrodes are disposed on the same plane defined by the substrate. In particular, in some of the embodiments described herein, a reagent is disposed proximate the at least two other electrodes and no reagent is disposed on the at least two electrodes. One feature of note in this system is the ability to provide for an accurate analyte measurement within about <NUM> seconds of deposition of a fluid sample (which may be a physiological sample) onto the biosensor as part of the test sequence.

As an example of an analyte calculation (e.g., glucose) for strip <NUM> (<FIG>, <FIG>, or <FIG> and its variants in FIGURES 3B-3T), it is assumed in <FIG> that the sampled signal value at <NUM> for the first working electrode <NUM> is about <NUM> nanoamperes whereas the signal value at <NUM> for the second working electrode <NUM> is about <NUM> nanoamperes and the calibration code of the test strip indicates that the Intercept is about <NUM> nanoamperes and the Slope is about <NUM> nanoamperes/mg/dL. Glucose concentration G<NUM> can be thereafter be determined from Equation <NUM> as follow: <MAT> where.

Intercept is the value obtained from calibration testing of a batch of test strips of which this particular strip comes from.

It is noted here that although the examples have been given in relation to a biosensor <NUM> which has two working electrodes (<NUM> and <NUM> in <FIG>) such that the measured currents from respective working electrodes have been added together to provide for a total measured current IE, the signal resulting from only one of the two working electrodes can be multiplied by two in a variation of test strip <NUM> where there is only one working electrode (either electrode <NUM> or <NUM>). Instead of a total signal, an average of the signal from each working electrode can be used as the total measured current IE for Equations <NUM>, <NUM>, and <NUM>-<NUM> described herein, and of course, with appropriate modification to the operational coefficients (as known to those skilled in the art) to account for a lower total measured current IE than as compared to an embodiment where the measured signals are added together. Alternatively, the average of the measured signals can be multiplied by two and used as IE in Equations <NUM>, <NUM>, and <NUM>-<NUM> without the necessity of deriving the operational coefficients as in the prior example. It is noted that the glucose concentration here is not corrected for any hematocrit value and that certain offsets may be provided to the signal values Iwe1 and Iwe2 to account for errors or delay time in the electrical circuit of the meter <NUM>. Temperature compensation can also be utilized to ensure that the results are calibrated to a referential temperature such as for example room temperature of about <NUM> degrees Celsius.

Now that a glucose concentration (G<NUM>) can be determined from the signal IE, a description of applicant's technique to determine hematocrit level of the fluid sample is provided in relation to <FIG>. In <FIG>, the system <NUM> (<FIG>) applies a first oscillating input signal <NUM> at a first frequency (e.g., of about 25kilo-Hertz) to a pair of sensing electrodes. The system is also set up to measure or detect a first oscillating output signal <NUM> from the third and fourth electrodes, which in particular involve measuring a first time differential Δt<NUM> between the first input and output oscillating signals. At the same time or during overlapping time durations, the system may also apply a second oscillating input signal (not shown for brevity) at a second frequency (e.g., about 100kilo-Hertz to about 1MegaHertz or higher, and preferably about <NUM> kilo Hertz) to a pair of electrodes and then measure or detect a second oscillating output signal from the third and fourth electrodes, which may involve measuring a second time differential Δt<NUM> (not shown) between the first input and output oscillating signals. From these signals, the system estimates a hematocrit level of the fluid sample based on the first and second time differentials Δt<NUM> and Δt<NUM>. Thereafter, the system is able to derive a glucose concentration. The estimate of the a hematocrit can be done by applying an equation of the form <MAT> where.

Details of this exemplary technique can be found in Provisional <CIT>, entitled, "Hematocrit Corrected Glucose Measurements for Electrochemical Test Strip Using Time Differential of the Signals" with Attorney Docket No. DDI-5124USPSP.

Another technique to determine a hematocrit level can be by two independent measurements of a hematocrit level. This can be obtained by determining: (a) the impedance of the fluid sample at a first frequency and (b) the phase angle of the fluid sample at a second frequency substantially higher than the first frequency. In this technique, the fluid sample is modeled as a circuit having unknown reactance and unknown resistance. With this model, an impedance (as signified by notation " | Z | ") for measurement (a) can be determined from the applied voltage, the voltage across a known resistor (e.g., the intrinsic strip resistance), and the voltage across the unknown impedance Vz; and similarly, for measurement (b) the phase angle can be measured from a time difference between the input and output signals by those skilled in the art. Details of this technique is shown and described in pending provisional patent application S. <CIT> (Attorney Docket No. DDI5215PSP). Other suitable techniques for determining the a hematocrit level of the fluid sample can also be utilized such as, for example, <CIT>, <CIT>, or <CIT>.

Another technique to determine the hematocrit can be obtained by knowing the phase difference (e.g., phase angle) and magnitude of the impedance of the sample. In one example, the following relationship is provided for the estimate of the physical characteristic or impedance characteristic of the sample ("IC"): <MAT> where:.

It is noted here that where the frequency of the input AC signal is high (e.g., greater than <NUM>) then the parametric terms y<NUM> and y<NUM> relating to the magnitude of impedance M may be ±<NUM>% of the exemplary values given herein such that each of the parametric terms may include zero or even a negative value. On the other hand, where the frequency of the AC signal is low (e.g., less than <NUM>), the parametric terms y<NUM> and y<NUM> relating to the phase angle P may be ±<NUM>% of the exemplary values given herein such that each of the parametric terms may include zero or even a negative value. It is noted here that a magnitude of H or HCT, as used herein, is generally equal to the magnitude of IC. In one exemplary implementation, H or HCT is equal to IC as H or HCT is used herein this application.

In another alternative implementation, Equation <NUM> is provided. Equation <NUM> is the exact derivation of the quadratic relationship, without using phase angles as in Equation <NUM>. <MAT> where:.

By virtue of the various components, systems and insights provided herein, a technique to achieve a temperature compensated analyte measurement can be understood with reference to <FIG>. This technique involves depositing a fluid sample (which may be a physiological sample or a control solution sample) on a biosensor at step <NUM> (e.g., in the form of a test strip as show in <FIG>, <FIG>, <FIG>) that has been inserted into a meter (step <NUM>). Once the meter <NUM> is turned on, a signal is applied to the strip <NUM> (or its variants) and when the sample is deposited onto the test chamber, the applied signal physically transforms the analyte (e.g., glucose) in the sample into a different physical form (e.g., gluconic acid) due to the enzymatic reaction of the analyte with the reagent in the test chamber. As the sample flows into the capillary channel of the test cell, at least one physical characteristic of the sample is obtained from an output of another signal driven into the sample (step <NUM>) along with estimate of the analyte concentration (step <NUM>). From the obtained physical characteristic (step <NUM>) and estimated analyte concentration (step <NUM>), a sampling time point is defined (at step <NUM>) at which the signal output from the sample during the test sequence is measured (at step <NUM>) and used for calculating the analyte concentration in step <NUM>. In particular, the step of obtaining the physical characteristic (step <NUM>) may include applying a first signal to the sample to measure a physical characteristic of the sample, while the step <NUM> of initiating an enzymatic reaction may involve driving a second signal to the sample, and the step of measuring (step <NUM>) may entail evaluating an output signal from the at least two electrodes at a point in time after the start of the test sequence, in which the point in time is set (at step <NUM>) as a function of at least the measured or estimated physical characteristic (step <NUM>) and estimated analyte concentration (step <NUM>).

The determination of the appropriate point (or time interval) during the test sequence TS as a function of the measured or estimated physical characteristic(s) (in step <NUM>) can be determined by the use of a look-up table programmed into the microprocessor of the system. For example, a look-up table may be provided that allows for the system to select the appropriate sampling time for the analyte (e.g., glucose or ketone) with measured or known physical characteristic comprising hematocrit and optionally viscosity of the sample.

In particular, an appropriate sampling time point may be based on an early estimation of the analyte and the measured or known physical characteristic to arrive at the appropriate sampling time that gives the lowest error or bias as compared to referential values. In this technique, a look up table is provided in which the defined sampling time point is correlated to (a) the estimated analyte concentration and (b) the physical characteristic of the sample. For example, Table <NUM> may be programmed into the meter to provide a matrix in which qualitative categories (low, medium, and high glucose) of the estimated analyte form the main column and the qualitative categories (low, medium, and high) of the measured or estimated physical characteristic form the header row. In the second column, t/Hct is a value determined experimentally of the time shift per % hematocrit difference from nominal hematocrit of <NUM>%. As one example, for <NUM>% hematocrit at "Mid-Glucose" would indicate a time shift of (<NUM> - <NUM>)*<NUM> = -<NUM>. The time of -<NUM> milliseconds is added to the original test time of about <NUM> milliseconds giving (<NUM>-<NUM>=<NUM> milliseconds) ~ <NUM>.

The time Tss (i.e., a specified sampling time) at which the system should be sampling or measuring the output signal of the biosensor is based on both the qualitative category of the estimated analyte and measured or estimated physical characteristic and is predetermined based on regression analysis of a large sample size of actual physiological fluid samples. Applicants note that the appropriate sampling time is measured from the start of the test sequence but any appropriate datum may be utilized in order to determine when to sample the output signal. As a practical matter, the system can be programmed to sample the output signal at an appropriate time sampling interval during the entire test sequence such as for example, one sampling every <NUM> milliseconds or even as little as about <NUM> milliseconds. By sampling the entire signal output transient during the test sequence, the system can perform all of the needed calculations near the end of the test sequence rather than attempting to synchronize the sampling time with the set time point, which may introduce timing errors due to system delay.

Applicant hereafter will discuss the look-up Table <NUM> in relation to the particular analyte of glucose in physiological fluid samples. Qualitative categories of blood glucose are defined in the first column of Table <NUM> in which low blood glucose concentrations of less than about <NUM>/dL are designated as "Lo-Glucose"; blood glucose concentrations of higher than about <NUM>/dL but less than about <NUM>/dL are designated as "Mid-Glucose"; and blood glucose concentrations of higher than about <NUM>/dL are designated as "Hi-Glucose".

During a test sequence, an "Estimated Analyte" can be obtained by sampling the signal at a convenient time point, typically at five seconds during a typical <NUM> seconds test sequence. The measurement sampled at this five second time point allows for an accurate estimate of the analyte (in this case blood glucose). The system may then refer to a look-up table (e.g., Table <NUM>) to determine when to measure the signal output from the test chamber at a specified sampling time Tss based on two criteria: (a) estimated analyte and (b) qualitative value of the physical characteristic of the sample. For criteria (b), the qualitative value of the physical characteristic is broken down into three sub-categories of Low Hct, Mid Hct and High Hct. Thus, in the event that the measured or estimated hematocrit level is high (e.g., greater than <NUM>%) and the estimated glucose is also high, then according to Table <NUM>, the test time for the system to measure the signal output of test chamber would be about <NUM> seconds. On the other hand, if the measured hematocrit is low (e.g., less than <NUM>%) and the estimated glucose is low then according to Table <NUM>, the test time Tss for the system to measure the signal output of test chamber would be about <NUM> seconds.

Once the signal output IT of the test chamber is measured at the designated time (which is governed by the measured or estimated physical characteristic), the signal IT is thereafter used in the calculation of the analyte concentration (in this case glucose) with Equation <NUM> below. <MAT> where.

It should be noted that the step of applying the first signal and the driving of the second signal is sequential in that the order may be the first signal then the second signal or both signals overlapping in sequence; alternatively, the second signal first then the first signal or both signals overlapping in sequence. Alternatively, the applying of the first signal and the driving of the second signal may take place simultaneously.

In the method, the step of applying of the first signal involves directing an alternating signal provided by an appropriate power source (e.g., the meter <NUM>) to the sample so that a physical characteristic of the sample is determined from an output of the alternating signal. The physical characteristic being detected is hematocrit. The directing step may include driving first and second alternating signal at different respective frequencies in which a first frequency is lower than the second frequency. Preferably, the first frequency is at least one order of magnitude lower than the second frequency. As an example, the first frequency may be any frequency in the range of about <NUM> to about <NUM> and the second frequency may be from about <NUM> to about <NUM> or more. As used herein, the phrase "alternating signal" or "oscillating signal" can have some portions of the signal alternating in polarity or all alternating current signal or an alternating current with a direct current offset or even a multi-directional signal combined with a direct-current signal.

Further refinements of Table <NUM> based on additional investigations of the technique allowed applicants to devise Table <NUM>, shown below.

As in Table <NUM>, a measured or estimated physical characteristic is used in Table <NUM> along with an estimated analyte concentration to derive a time Tss at which the sample is to be measured. For example, if the measured charactertistic is about <NUM>% and the estimated glucose (e.g., by sampling at about <NUM> to <NUM> seconds) is about <NUM>, the time at which the microcontroller should sample the fluid is about <NUM> seconds. In another example, where the estimated glucose is about <NUM>/dL and the measured or estimated physical characteristic is <NUM>%, the specified sampling time would be about <NUM> seconds.

For the embodiments utilized with Table <NUM>, the estimated glucose concentration is provided with an equation: <MAT> where.

From the estimated glucose, the glucose concentration can be determined from: <MAT> where:.

Although applicant's technique may specify only one sampling time point, the method may include sampling as many time points as required, such as, for example, sampling the signal output continuously (e.g., at specified sampling time such as, every <NUM> milliseconds to <NUM> milliseconds) from the start of the test sequence until at least about <NUM> seconds after the start and the results stored for processing near the end of the test sequence. In this variation, the sampled signal output at the specified sampling time (which may be different from the predetermined sampling time point) is the value used to calculate the analyte concentration.

It is noted that in the preferred embodiments, the measurement of a signal output for the value that is somewhat proportional to analyte ( glucose) concentration is performed prior to the estimation of the hematocrit. Alternatively, the hematocrit level can be estimated prior to the measurement of the preliminary glucose concentration. In either case, the estimated glucose measurement GE is obtained by Equation <NUM> with IE sampled at about one of <NUM> seconds or <NUM> seconds, as in <FIG>, the Hct is obtained by Equation <NUM> and the glucose measurement G is obtained by using the measured signal output ID at the designated sampling time point(s) (e.g., the measured signal output ID being sampled at <NUM> seconds or <NUM> seconds) for the signal transient <NUM>.

Other techniques for determining the analyte concentration or value are shown and described in <CIT>, <CIT>; <CIT>.

Referring back to step <NUM> in <FIG>, the system evaluates whether a value defined by a difference in the magnitudes of the respective signal outputs of the first and second electrodes (Iwe1 and Iwe2) divided by the magnitude of the signal output of the second electrode is greater than a predetermined threshold Pth. Applicant has utilized the evaluation step <NUM> for the system due to the summation of the output signals from the first and second working electrodes. Because both electrodes are configured to undergo similar electrochemical reactions, both electrodes should have the same magnitude for their respective signal outputs. Referring to <FIG>, the output signals from respective working electrodes are shown as being virtually identical for the entire time interval at which the output signals from the working electrodes are sampled. However, when there is insufficient fluid sample volume or other environmental factors (e.g., humidity and temperature), both electrodes may not undergo similar electrochemical reactions, thereby skewing at least one of the output signals and leading to an incorrect analyte result being annunciated to the user at step <NUM>. This less than ideal condition can be seen in <FIG> where there is a clear divergence "▲" in the magnitudes of the output signals from respective working electrodes due to one or more such electrodes failing to receive sufficient sample volume or a defect in the enzyme layer disposed on the electrode. Regardless of the reasons, under such a condition shown in <FIG>, the sum of the magnitudes of the signals from the first and second working electrodes (Iwe1 and Iwe2) may provide an incorrect analyte concentration.

Consequently, applicant has devised a solution to this problem of determining when to annunciate that there is an error in the filling of the fluid sample. In particular, applicant has devised a test in which the output signals from both electrodes are compared using a bias of the two electrodes to each other and compared to a predetermined threshold.

The predetermined threshold can be from about <NUM> to about <NUM> and preferably about <NUM>. The mathematical representation of the evaluation that would trigger an error is shown by Eq. <NUM>: <MAT>.

In the evaluation at step <NUM>, if the value (e.g., (Iwe1 - Iwe2)*<NUM>/Iwe2) is greater than the predetermined threshold Pth then the system would annunciate an error (step <NUM>) and terminate further processing (step <NUM>). On the other hand, if the value is less than the predetermined threshold Pth then the system may proceed to step <NUM> to determine or calculate an analyte concentration from the sampled output signals of respective first and second electrodes at the specified time point. At step <NUM>, the system may annunciate the analyte concentration determined by the system.

Applicant notes that the technique is designed so that if such fill error is detected, the system will quickly annunciate an error (from step <NUM> directly to step <NUM>) and terminate the assay process.

An alternative technique has also been devised that allows for the system to set an error flag while allowing the continuation of the acquisition of an analyte concentration and then terminating the assay only thereafter. In particular, this technique can be achieved with reference to step <NUM> where it is assumed that the fill error value is greater than a preset threshold such that the process moves to step <NUM> to set a fill error flag as active. Thereafter, the process moves to step <NUM> to continue with the acquisition of the analyte concentration. It is only after step <NUM> that the system may query to see if one or more error flags (including the fill error flag) have been set. If a certain number of error flags (including a minimum of just one fill error flag) have been set, the system will immediately annunciate the error at step <NUM> and thereafter terminating the assaying process at step <NUM>.

As is known, the detection of the hematocrit does not have to be done by alternating signals but can be done with other techniques.

As described earlier, the microcontroller or an equivalent microprocessor (and associated components that allow the microcontroller to function for its intended purpose in the intended environment such as, for example, the processor <NUM> in <FIG>) can be utilized with computer codes or software instructions to carry out the methods and techniques described herein. Applicants note that the exemplary microcontroller <NUM> (along with suitable components for functional operation of the processor <NUM>) in <FIG> is embedded with firmware or loaded with computer software representative of the logic diagrams in <FIG> and the microcontroller <NUM>, along with associated connector <NUM> and interface <NUM> and equivalents thereof, are the means for: (a) determining a specified sampling time based on a sensed or estimated physical characteristic, the specified sampling time being at least one time point or interval referenced from a start of a test sequence upon deposition of a sample on the test strip and (b) determining an analyte concentration based on the specified sampling time. Alternatively, the means for determining may include means for applying a first signal to the plurality of electrodes so that a batch slope defined by a physical characteristic of a fluid sample is derived and for applying a second signal to the plurality of electrodes so that an analyte concentration is determined based on the derived batch slope and the specified sampling time. Furthermore, the means for determining may include means for estimating an analyte concentration based on a predetermined sampling time point from the start of the test sequence and for selecting a specified sampling time from a matrix of estimated analyte concentration and sensed or estimated physical characteristic. Yet further, the means for determining may include means for selecting a batch slope based on the sensed or estimated physical characteristic and for ascertaining the specified sampling time from the batch slope.

Claim 1:
An analyte measurement system comprising:
a test strip (<NUM>) including:
a substrate (<NUM>);
a plurality of electrodes connected to respective electrode connectors;
a sample chamber (<NUM>) comprising the plurality of electrodes and an enzymatic reagent (<NUM>); and
an analyte meter (<NUM>) including:
a housing;
a test strip port connector (<NUM>) configured to connect to the respective electrode connectors of the test strip (<NUM>); and
a microprocessor (<NUM>) in electrical communication with the test strip port connector (<NUM>) to apply electrical signals or sense electrical signals from the plurality of electrodes, wherein the microprocessor (<NUM>) is configured to:
(a) apply a second signal to a first electrode and a second electrode of the plurality of electrodes to initiate an enzyme reaction of an analyte with the reagent;
(b) apply a first signal to the plurality of electrodes and determine a physical characteristic of a fluid sample (<NUM>) from the output resulting from the first signal, wherein the physical characteristic comprises a blood hematocrit level;
(c) estimate an analyte concentration (Gest) from the output resulting from the first signal at a predetermined sampling time point during a test sequence, wherein the estimated analyte concentration (Gest) comprises a blood glucose concentration;
(d) determine a specified sampling time point (Tss) from the determined physical characteristic and the estimated analyte concentration (Gest);
(e) measure a signal output (Iwe1, Iwe2) at the specified sampling time point (Tss) from each of the first and second electrodes;
(f) evaluate whether a value defined by a difference in the magnitudes of the respective signal outputs of the first and second electrodes (Iwe1, Iwe2) divided by the magnitude of the signal output of the second electrode (Iwe2) is greater than a predetermined threshold (Pth);
(g) if the value is less than the predetermined threshold (Pth) then determine or calculate the analyte concentration (Go) from the signal outputs of the first and second electrodes (Iwe1, Iwe2) at the specified sampling time (Tss) and annunciate the analyte concentration, wherein the analyte concentration comprises a blood glucose concentration; and
(h) if the value is greater than the predetermined threshold (Pth) then annunciate an error.