Method and apparatus for analyte detection using electrochemical biosensor

Abstract

To sense an analyte utilizing a sensor having a working electrode.SOLUTION: A method includes providing the working electrode with an analyte-specific enzyme and a redox mediator, providing the working electrode to the analyte, accumulating charge derived from the analyte reacting with the analyte-specific enzyme and the redox mediator for a prescribed period of time, connecting the working electrode to a circuit after the prescribed period of time, and measuring the signal from the accumulated charge.SELECTED DRAWING: None

Description

[Technical field]

[0001] This patent application claims the benefit of priority to U.S. Provisional Application No. 16 / 024,353, filed June 29, 2018; U.S. Provisional Application No. 62 / 527,981, filed June 30, 2017; U.S. Provisional Application No. 62 / 544,692, filed August 11, 2017; and U.S. Provisional Application No. 62 / 545,252, filed August 14, 2017, the contents of which are incorporated herein by reference in their entireties.

[0002] This invention was made with Government support under Contract No. HDTRA-1-16-C-0048 awarded by the Defense Threat Reduction Agency. The Government has certain rights in this invention.

[0003] Some aspects of the present disclosure relate to sensing of an analyte using an electrochemical enzyme biosensor. For example, some aspects of the present disclosure relate to methods and enzyme biosensors that allow for the detection of low concentrations of an analyte by accumulating the analyte on the biosensor. [Background technology]

[0004] Enzyme biosensors, which use an enzyme in combination with a transducer that serves as a biorecognition element for the target analyte, have been developed and used. A variety of different signal transduction methods have been used, but the most frequently used is an electrochemical signal. Electrochemical biosensors can directly convert a biological event (such as analyte detection) into an electrical signal, so they do not require complex equipment, and as a result, electrochemical biosensors have desirable characteristics in terms of size, cost, and portability. Among the electrochemical techniques used for signal transduction, amperometry is often used. Amperometry is a method in which the current flowing through the sensor is measured while the working electrode of the sensor is maintained at a constant potential (voltage). Such sensors are designed to pass a current that is dependent on the analyte concentration.

[0005] An example of an enzymatic biosensor using amperometry is the continuous glucose sensor, a wearable in vivo device designed to allow the user to measure blood glucose levels frequently. These devices use glucose oxidase (GO) immobilized on the working electrode as the glucose sensing element. x Amperometry utilizes a glucose oxidoreductase enzyme such as .Electrons are first transferred from glucose to the enzyme via enzymatic oxidation and then transferred to the working electrode via a redox mediator such as oxygen (O2) or an osmium (Os)-containing redox polymer. Although amperometry has proven effective in measuring analytes such as glucose that are present in relatively high physiological concentrations (5 millimolar (mM) or greater), it may not be suitable for measuring analytes that are present in low concentrations. Summary of the Invention [Means for solving the problem]

[0006] Aspects of embodiments of the present disclosure relate to methods for detecting low concentrations of an analyte (eg, less than 5 mM, 1 nanomolar (nM) to 5 mM, or 4.7 nM to 5 mM) by accumulating the analyte in an enzyme biosensor.

[0007] According to some aspects of the present disclosure, there is provided a method of sensing an analyte using a sensor having a working electrode, the method comprising providing an analyte-specific enzyme and a redox mediator to the working electrode, providing the working electrode to the analyte, accumulating charge from the analyte reacting with the analyte-specific enzyme and the redox mediator over a period of time, connecting the working electrode to a circuit after the period of time, and measuring a signal from the accumulated charge.

[0008] According to some aspects of the present disclosure, the method includes connecting the working electrode to a circuit prior to providing the working electrode to the analyte, and disconnecting the working electrode from the circuit prior to providing the working electrode to the analyte.

[0009] According to some aspects of the present disclosure, the method includes connecting the working electrode to a circuit prior to providing the working electrode to the analyte, and disconnecting the working electrode from the circuit prior to providing the working electrode to the analyte.

[0010] According to some aspects of the present disclosure, the sensor is an enzymatic electrochemical biosensor.

[0011] According to some aspects of the present disclosure, the redox mediator is an immobilized redox polymer.

[0012] According to some aspects of the present disclosure, the immobilized redox polymer comprises a redox species and a polymer, the redox species is selected from an osmium (Os), ruthenium (Ru), iron (Fe), or cobalt (Co) containing polymer, and the polymer is selected from poly(vinylpyridine), poly(thiophene), poly(aniline), poly(pyrrole), or poly(acetylene).

[0013] According to some embodiments of the present disclosure, the immobilized redox polymer is an Os-containing poly(vinylpyridine).

[0014] According to some aspects of the present disclosure, the analyte is selected from cortisol, glucose, lactate, 3-hydroxybutyrate, alcohol, pyruvate, glutamate, theophylline, or creatinine.

[0015] According to some aspects of the present disclosure, the analyte-specific enzyme is a nicotinamide adenine dinucleotide (NAD)-dependent dehydrogenase, a flavin adenine dinucleotide (FAD)-dependent oxidase, and / or a flavin mononucleotide (FMN)-dependent oxidase.

[0016] According to some embodiments of the present disclosure, the analyte-specific enzyme is selected from 11β-hydroxysteroid dehydrogenase type 2 (11β-HSD-2), glucose oxidase, NAD-glucose dehydrogenase, FAD-glucose dehydrogenase, lactate oxidase, NAD-lactate dehydrogenase, NAD-alcohol dehydrogenase, pyruvate oxidase, NAD-glutamate dehydrogenase, or xanthine oxidase.

[0017] According to some aspects of the present disclosure, the accumulation of charge includes accumulation of electrons.

[0018] According to some aspects of the present disclosure, the sensor is placed subcutaneously in the subject.

[0019] According to some embodiments of the present disclosure, the analyte is at a concentration as low as 4.7 nanomolar (nM).

[0020] According to some aspects of the present disclosure, the predetermined period of time ranges from 60 seconds to 30 minutes. According to some aspects, the predetermined period of time ranges from 120 seconds to 30 minutes. According to some aspects, the predetermined period of time ranges from 120 seconds to 10 minutes.

[0021] According to some aspects of the present disclosure, the sensor includes an outer membrane. According to some aspects, the outer membrane is a flow restricting membrane. According to some aspects, the outer membrane is an analyte permeable membrane.

[0022] According to some aspects of the present disclosure, measuring the signal from the accumulated charge includes measuring a peak height of the signal and / or measuring a peak area of ​​the signal.

[0023] According to some aspects, the method further comprises calibrating the measured peak height to provide a concentration of the analyte.

[0024] According to some aspects, the method further comprises calibrating the measured peak area to provide a concentration of the analyte.

[0025] According to some embodiments, measuring the signal from the accumulated charge includes recording the signal at a sampling rate of 0.1 to 0.5 Hertz (Hz) and / or filtering the signal at a frequency of 0.032 to 3.2 Hertz (Hz).

[0026] According to some aspects of the present disclosure, the working electrode includes a sensing element comprising the analyte-specific enzyme and the redox mediator. According to some aspects, the sensing element comprises carbon nanotubes.

[0027] According to some aspects, there is provided a method of sensing an analyte utilizing a sensor comprising a working electrode comprising an analyte specific enzyme and a redox mediator, the method comprising providing the working electrode to the analyte, accumulating charge from the analyte reacting with the analyte specific enzyme and the redox mediator, and measuring a signal from the accumulated charge by measuring a peak height of the signal and / or measuring a peak area of ​​the signal.

[0028] According to some aspects of the present disclosure, there is provided a system for sensing an analyte, comprising a working electrode, a sensing element, and a circuit, the sensing element comprising an analyte-specific enzyme and a redox mediator, the sensing element being disposed on the working electrode and configured to accumulate charge derived from the analyte reacting with the analyte-specific enzyme and the redox mediator over a period of time, the circuit being connected to the working electrode after a period of time and configured to measure a signal from the accumulated charge. According to some aspects, the sensing element of the system comprises a carbon nanotube. According to some aspects, the system comprises an outer membrane covering at least the sensing element. According to some aspects, the analyte-specific enzyme of the system is selected from nicotinamide adenine dinucleotide (NAD)-dependent dehydrogenase, flavin adenine dinucleotide (FAD)-dependent oxidase, or flavin mononucleotide (FMN)-dependent oxidase. For example, in some embodiments, the analyte-specific enzyme of the system is selected from 11β-hydroxysteroid dehydrogenase type 2 (11β-HSD-2), glucose oxidase, NAD-glucose dehydrogenase, FAD-glucose dehydrogenase, lactate oxidase, NAD-lactate dehydrogenase, NAD-alcohol dehydrogenase, pyruvate oxidase, NAD-glutamate dehydrogenase, and xanthine oxidase. [Brief description of the drawings]

[0029] [Figure 1] FIG. 1 is a flow diagram illustrating a storage mode sensing method including acts 10, 15, 20, 25, and 30 according to an embodiment of the present disclosure.

[0030] [Diagram 2]2 is a schematic diagram showing an electrode setup used for accumulation mode sensing according to an embodiment of the present disclosure. When the circuit is connected as shown in the left frame, the working electrode is at a potential (voltage) sufficient to drive the redox reaction of the analyte under steady state conditions, and when the circuit is broken as shown in the right frame, the working electrode is electrically disconnected from the circuit allowing electrons from the analyte to be stored in the redox polymer until the working electrode is reconnected with the circuit and the stored charge can be measured.

[0031] [Figure 3A] FIG. 3A shows predicted current versus time signals and certain quantitative parameters (accumulation time, peak area, and peak height when the circuit is disconnected as shown) for sensing in accumulation mode according to an embodiment of the present disclosure.

[0032] [Figure 3B] FIG. 3B shows a schematic of the redox reaction occurring during accumulation mode sensing (circuit shown as “break circuit” as shown) of an oxidizable analyte (analyte A) using an oxidase enzyme (AOx) co-immobilized with an osmium redox polymer (Os3+) in accordance with an embodiment of the present disclosure.

[0033] [Figure 3C] FIG. 3C shows current versus time waveforms obtained for accumulation mode sensing of 2 μM glucose (shown in white) according to an embodiment of the present disclosure, measured at five different accumulation times using an example glucose sensor (at +40 mV as shown by the diagonal lines).

[0034] [Figure 3D] FIG. 3D shows a calibration curve of the accumulation mode signal measured by peak height or peak area versus current measurement and accumulation time shown in FIG. 3C according to an embodiment of the present disclosure.

[0035] [Figure 4A]FIG. 4A shows a representative current versus time waveform for a calibration experiment using accumulation mode sensing with an example glucose sensor (at +40 mV as shown by the hatched line) with a 60 second accumulation time for each detection (when the circuit is shown in white) in accordance with an embodiment of the present disclosure.

[0036] [Figure 4B] FIG. 4B shows a comparison of current measurements and a calibration curve obtained from accumulation mode signals measured in a sensing experiment according to the embodiment of the present disclosure shown in FIG. 4A.

[0037] [Diagram 5] 5 shows calibration curves for amperometric and accumulation mode sensing (peak height and peak area) measurements at glucose concentrations of 0, 50, 100, 200, and 500 μM with accumulation times of 1 minute (diamonds), 2 minutes (triangles), 5 minutes (squares), and 10 minutes (circles) according to embodiments of the present disclosure, as indicated. Each calibration curve shows the average response of four sensors.

[0038] [Figure 6A] FIG. 6A shows a graph of potential versus time signal for an exemplary glucose sensor obtained using the open circuit potential method sensing various nanomolar (nM) concentrations of glucose as shown, in accordance with an embodiment of the present disclosure.

[0039] [Figure 6B] FIG. 6B shows a calibration curve (slope versus glucose concentration (nM)) for the data in the graph of FIG. 6A according to an embodiment of the present disclosure.

[0040] [Figure 6C] FIG. 6C shows a graph of potential versus time signal for an exemplary glucose sensor obtained using the open circuit potential method for detecting glucose at various nM concentrations as shown, in accordance with an embodiment of the present disclosure.

[0041] [Figure 6D]FIG. 6D shows the slope of the calibration curve (glucose concentration (nM)) for the data of the graph of FIG. 6C according to an embodiment of the present disclosure.

[0042] [Figure 6E] FIG. 6E shows a composite calibration curve of an exemplary glucose sensor (filled circle data points, n=8) and a control sensor (open circle data points, n=4) from in vitro detection of glucose using the open circuit potential method according to an embodiment of the present disclosure.

[0043] [Figure 6F] FIG. 6F shows a zoomed-in view of the calibration curve of FIG. 6E obtained from 0-200 nM glucose according to an embodiment of the present disclosure.

[0044] [Figure 6G] FIG. 6G shows a graph of the potential versus time signal of an exemplary glucose sensor obtained using an open circuit potential method using an exemplary glucose sensor sensing various nM concentrations of glucose as shown as the working electrode and a control sensor (containing redox polymer but no glucose oxidase) as the reference electrode according to an embodiment of the present disclosure.

[0045] [Figure 6H] FIG. 6H shows the slope of the calibration curve (glucose concentration (nM)) for the data of the graph of FIG. 6G according to an embodiment of the present disclosure.

[0046] [Figure 7] FIG. 7 shows a comparison of accumulation mode signal shapes at different filtering frequencies, namely 3.2 Hz (shown as a solid black line) and 0.032 Hz (shown as a dashed line), in accordance with an embodiment of the present disclosure.

[0047] [Figure 8A] FIG. 8A shows micrographs of deposited glucose sensing reagent according to an embodiment of the present disclosure with (right frame) and without (left frame) carbon nanotubes (CNTs).

[0048] [Figure 8B] FIG. 8B shows calibration curves of amperometric and accumulation mode detection (peak height and peak area) using different filtering frequencies (0.032 Hz, circle marks, and 3.2 Hz, triangle marks) and sensing reagents with and without CNTs in accordance with an embodiment of the present disclosure.

[0049] [Figure 9A] 9A shows accumulation mode signals obtained with a representative glucose sensor during a calibration experiment using glucose concentrations from 0 to 200 nM, in accordance with an embodiment of the present disclosure, where the accumulation time for each detection was 30 min, the signal was filtered at 3.2 Hz, and CNTs were added to the sensing reagent.

[0050] [Figure 9B] 9B shows the corresponding linear regression calibration curves obtained from the amperometric and accumulation mode signals measured in the sensing experiment shown in FIG. 8A, according to an embodiment of the present disclosure. Each signal is the background-subtracted average of eight sensors, and the error bars indicate the standard deviation. The bottom row of the plots is a zoomed in view of glucose concentrations from 0 to 50 nM.

[0051] [Figure 10A] FIG. 10A shows accumulation mode signals from a representative glucose sensor under background conditions ([glucose]=0), in air (bold line), and in oxygen-purged buffer (thin line) in accordance with an embodiment of the present disclosure.

[0052] [Figure 10B] 10B shows a schematic of the background current measurement and accumulation mode signal from the experiment shown in FIG 10A, where the signal is an average of four sensors, with oxygen purged data shown as solid circles and air data shown as open circles, according to an embodiment of the present disclosure.

[0053] [Figure 11]11 shows calibration curves obtained for amperometric and accumulation mode sensing (peak height and peak area) during sensing experiments performed at glucose concentrations from 0 to 200 μM according to an embodiment of the present disclosure, where each line is a linear best fit line obtained at concentrations from 0 to 200 nM, extrapolated to higher concentrations, and each signal is an average of eight sensors.

[0054] [Figure 12] FIG. 12 shows a schematic diagram of an analyte sensor according to an embodiment of the present disclosure.

[0055] [Figure 13] FIG. 13 is a cross-sectional view of a portion of an analyte sensor consistent with one or more embodiments of the present disclosure.

[0056] [Figure 14A] FIG. 14A shows a plan view of an implantable analyte sensor consistent with one or more embodiments of the present disclosure.

[0057] [Figure 14B] FIG. 14B is a cross-sectional view of a portion of any analyte sensor having a membrane consistent with one or more embodiments of the present disclosure.

[0058] [Figure 14C] FIG. 14C shows a close-up view of the outer membrane-covered sensing layer, working electrode, and substrate according to an embodiment of the present disclosure.

[0059] [Figure 14D] FIG. 14D is a schematic diagram illustrating a redox reaction of an analyte with an analyte-specific enzyme and a redox mediator at a working electrode according to an embodiment of the present disclosure.

[0060] [Figure 15] FIG. 15 is a block diagram of one embodiment of an analyte monitoring system in accordance with an embodiment of the present disclosure.

[0061] [Figure 16]FIG. 16 is a block diagram of one embodiment of a reader of the analyte monitoring system of FIG. 15 in accordance with an embodiment of the present disclosure.

[0062] [Figure 17] FIG. 17 is a block diagram of one embodiment of a sensor control unit of the analyte monitoring system of FIG. 15 in accordance with an embodiment of the present disclosure. Detailed Description of the Invention

[0063] Embodiments of the present disclosure provide electrochemical measurement methods using electrochemical sensors to measure low nanomolar concentrations of analytes in vitro and in vivo. Embodiments of the present disclosure include electrochemical sensors, such as enzyme biosensors, that have been modified to measure low nanomolar concentrations of analytes.

[0064] Where a numerical range is provided, to the tenth of the unit of the lower limit unless the context clearly indicates otherwise, each value between the upper and lower limits of that range is also specifically disclosed. Each smaller range included between any stated or intervening value in a stated range, and any other stated or intervening value in that stated range, is encompassed within this disclosure. The upper and lower limits of these smaller ranges are independently included or excluded within the stated range. Each range in which either or both of the upper and lower limits are not included in the narrower range, or both are included, is also encompassed within this disclosure, as are any specifically excluded limits within the stated range. Where a stated range includes one or both of the upper and lower limits, ranges excluding either or both of those included upper and lower limits are also included within this disclosure.

[0065] As used herein, the terms "substantially," "about," and similar terms are used as terms of approximation rather than degree and are intended to account for inherent deviations in measurements or calculations that would be apparent to one of ordinary skill in the art.

[0066] As will be appreciated, in the description disclosed herein, unless otherwise indicated, either implicitly or explicitly, singular words include the plural and plural words include the singular. By way of example only, unless otherwise indicated, either implicitly or explicitly, a reference to an "analyte" with "an" or "the" includes a combination and / or mixture of two or more different analytes, in addition to a single analyte. Also, a reference to a "concentration value" with "an" or "the" includes two or more concentration values, in addition to a single concentration value. Furthermore, as will be appreciated, for any given component described herein, generally, any of the possible candidates or options listed for that component may be used individually or in combination with each other, unless otherwise indicated, either implicitly or explicitly. Also, as will be appreciated, any list of such candidates or options is merely illustrative and not limiting, unless otherwise indicated, either implicitly or explicitly.

[0067] As used herein, the terms "measure," "measurement," and "measured" may encompass one meaning each of the terms "determine," "determined," "determined," "calculate," "calculate," and "calculated."

[0068] As used herein, an "electrochemical sensor" is a device adapted to detect the presence and / or measure the level of an analyte in a sample via electrochemical redox reactions at the sensor surface. These reactions are converted into an electrical signal that may be correlated to the amount, concentration, or level of the analyte in the sample.

[0069] As used herein, a "working electrode" is an electrode at which an analyte (or a second compound whose level depends on the level of the analyte) is electrooxidized or electroreduced, with or without the action of an electron transfer agent.

[0070] As used herein, a "counter electrode" refers to an electrode that is paired with a working electrode. A counter electrode passes a current that is equal in magnitude and opposite in sign to the current passing through the working electrode. In the context of the embodiments of the present disclosure, unless otherwise specified, the term "counter electrode" includes both a) a counter electrode and b) a counter electrode that also functions as a reference electrode (i.e., a counter electrode / reference electrode).

[0071] As used herein, unless otherwise specified, a "reference electrode" includes both a) a reference electrode and b) a reference electrode that also functions as a counter electrode (ie, a counter / reference electrode).

[0072] As used herein, "electrolysis" refers to the electro-oxidation or electro-reduction of a compound at an electrode, either directly or via one or more electron transfer agents.

[0073] As used herein, a component is "immobilized" within a sensor if, for example, the component is trapped within a component of the sensor or is covalently, ionically, or coordinatively bound to and / or trapped within a polymer, sol-gel matrix, or membrane that prevents migration.

[0074] As used herein, an "electron transfer agent" is a compound that carries electrons between the analyte and the working electrode, either directly or in combination with other electron transfer agents. One example of an electron transfer agent is a redox mediator.

[0075] As used herein, a "redox mediator" is an electron transfer agent that transfers electrons between an analyte, an enzyme that reduces or oxidizes the analyte, and an electrode, either directly or via one or more additional electron transfer agents. Redox mediators that include a polymeric backbone may also be referred to as "redox polymers."

[0076] As used herein, the term "precursor polymer" refers to the starting polymer before the various modifier groups are added to form the modified polymer.

[0077] As used herein, a "sensing layer" is a component of a sensor that includes components that facilitate electrolysis of an analyte. The sensing layer may include components such as an electron transfer agent (e.g., a redox mediator or redox polymer), a catalyst (e.g., an analyte-specific enzyme) that catalyzes a reaction of the analyte to generate a response at the working electrode, or both an electron transfer agent and a catalyst. In some embodiments of the present disclosure, the sensor includes a sensing layer disposed non-leachingly near or on the working electrode.

[0078] As used herein, a "sensing element" is a coating or area of ​​an analyte-specific enzyme disposed with a sensing layer. The sensing element is thus capable of interacting with an analyte. The sensing layer may have two or more sensing elements that form an analyte detection area disposed on the working electrode surface. In some embodiments, the sensing element comprises an analyte-specific enzyme and an electron transfer agent (e.g., a redox mediator). In some embodiments, the sensing element comprises an analyte-specific enzyme, an electron transfer agent, and a crosslinker.

[0079] As used herein, a "non-leachable" or "non-releasable" compound, or a compound "designed not to be leached," is meant to define a compound that is added to the sensor such that it does not substantially diffuse away from the sensing layer of the working electrode while the sensor is in use (e.g., while the sensor is implanted in a patient or while a sample is being measured).

[0080] As used herein, a "crosslinker" is a molecule that contains at least two reactive groups capable of linking at least two molecules or linking at least two portions of the same molecule. Linking at least two molecules is referred to as intermolecular crosslinking, and linking at least two portions of the same molecule is referred to as intramolecular crosslinking. Crosslinkers with three or more reactive groups are capable of both intermolecular and intramolecular crosslinking simultaneously.

[0081] A "membrane solution" is a solution containing all the ingredients necessary for crosslinking and membrane formation, including a modified polymer containing heterocyclic nitrogen groups, a crosslinking agent, and a buffer or an alcohol / buffer mixed solvent.

[0082] As used herein, "body fluid" or "biological fluid" refers to any body fluid or body fluid derivative in which an analyte may be measured, including, for example, blood, interstitial fluid, plasma, dermal fluid, sweat, and tears.

[0083] As used herein, "accumulation mode sensing" refers to the accumulation of electrons generated from the oxidation of an analyte, which occurs at or on a sensing element of a working electrode that is not connected to a circuit, thereby causing the accumulation of electrons.

[0084] Accumulation mode sensing Referring to the method flow diagram of FIG. 1, some embodiments of the present disclosure include a method of obtaining a signal from an analyte using a sensor including a working electrode and another electrode (e.g., a counter electrode and / or a reference electrode). The working electrode is provided with or modified with a catalyst (e.g., an analyte-specific enzyme) and an electron transfer agent (e.g., a redox mediator) (10). The area of ​​the working electrode modified with the analyte-specific enzyme and redox mediator may be referred to as the sensing element or sensing layer of the working electrode. As shown in FIG. 1, the working electrode provided with (e.g., modified with) the analyte-specific enzyme is provided with an analyte (15). In the presence of the analyte, the modified working electrode oxidizes the analyte and the amount of oxidation is measured as the amount of charge generated from the reaction. Charge from the redox reaction continues to accumulate on the working electrode surface (20) unless the working electrode is connected to another electrode. For analytes with low concentrations in the body (e.g., cortisol), the accumulation of charge (electrons) over a period of time provides a signal output from the low concentration of the analyte, which is easier to measure and quantify than other known methods. After accumulating charge for a period of time (e.g., up to 120 seconds, up to 3 minutes, up to 5 minutes, up to 10 minutes, up to 15 minutes, up to 20 minutes, up to 25 minutes, or up to 30 minutes), the working electrode is connected to at least one other electrode (e.g., a counter electrode and / or a reference electrode) to form a circuit (25). Once the circuit is formed, electrons accumulated on the working electrode surface are released as an electrical signal whose amplitude is measured (30) and correlated to the amount of analyte present at the working electrode. Thus, low concentrations (e.g., nanomolar amounts, such as 4.7 nM) of analyte can be readily detected and measured by the following method according to an embodiment of the present disclosure, shown as acts 10, 15, 20, 25, and 30 in FIG. 1.

[0085] Referring to FIG. 2, an example three electrode setup is shown with the working electrode 40, the reference electrode 50, and the counter electrode 60 used for accumulation mode sensing according to an embodiment of the present disclosure. In this figure, the circuit 70 is connected as shown in the left frame, and the working electrode is at a potential (voltage) sufficient to drive the redox reaction of the analyte under steady state conditions. For example, in the glucose sensor used herein, the potential (voltage) sufficient to drive the redox reaction is +40 mV vs. Ag / AgCl. When the circuit 70 is not connected as shown in the right frame, the working electrode 40 is electrically disconnected from the circuit 70, and charge (e.g., electrons) from the analyte can be stored in the redox polymer until the working electrode 40 is reconnected with the circuit 70 and the stored charge is measured.

[0086] 3A and 3B, an example of an electrochemical enzymatic biosensor is shown in a conceptual diagram in accumulation mode. In this example, the analyte (A) is reacted with an oxidoreductase (AO) that is electrically "connected" to the working electrode of the sensor via a redox polymer. x ) during normal amperometric sensing. During normal amperometric sensing, the electrodes are at a potential (voltage) such that the analyte reacts at a constant rate. The reaction rate is proportional to the analyte concentration. The oxidation reaction of the analyte (A = A + In the case of the ion exchange reaction (OECD), electrons are transported at a constant rate from the analyte (A) to the analyte-specific enzyme (AO), as shown in Figure 3B. x ), redox polymers (e.g., Os3 + ) to the working electrode, generating a steady-state current as shown in Figure 3A. When the working electrode is disconnected from the circuit, electrons stop flowing from the redox polymer to the working electrode, and no current flows through the circuit. However, the analyte is still subject to enzymatic oxidation, which results in the reduction of the redox polymer (Os3 + From Os2 + ) from the specimen (e - ) is stored in the redox polymer, which allows the reduced form of the redox polymer (Os2) to be converted over time. + ) stacking (Os2 +When the working electrode is reconnected to the circuit and brought to its original potential (voltage), the redox polymer stack is oxidized, resulting in a large current spike as shown in Figure 3A. The current then drops to the original amperometric current as the redox system reaches steady state again. This two-step process forms the basis of accumulation mode sensing. In the first step, the working electrode of the sensor is disconnected from the circuit for a period of time (also called the accumulation time), allowing charge from the analyte to be "stored" in the redox polymer. In the second step, after the accumulation time, the working electrode of the sensor is connected to the circuit, releasing the stored charge, which can be measured as a sharp peak.

[0087] 3C and 3D, an example of accumulation mode sensing is shown using the developed glucose sensor consisting of a glucose-specific sensing reagent deposited on a screen-printed carbon electrode. -Consisting of a glucose oxidase enzyme cross-linked to a redox polymer, this reagent has already been shown for use in glucose biofuel cells and in both self-powered and potentiostatically powered continuous glucose sensors. See, e.g., Mao et al., J. Am. Chem. Soc. 2003, 125:4951-4957; Mano et al., J. Am. Chem. Soc. 2003, 125:6588-6594; Liu et al., Anal. Chem. 2012, 84:3403-3409; Feldman et al., Diabetes Technol. Ther. 2003, 5:769-779; Hoss et al., J. Diabetes Sci. Technol. 2013, 7:1210-1219; and Hoss et al., J. Diabetes Sci. Technol. 2014, 8:89-94, the entire contents of which are incorporated herein by reference. In some embodiments of the present disclosure, the sensitivity of electrochemical measurements may be increased using an accumulation mode sensing method. In the experiments shown in Figures 3C and 3D, the glucose sensor was placed in a solution of 2 μM glucose and 100 mM phosphate buffered saline (PBS) and several accumulation mode measurements were made while monitoring the sensor current. In each measurement, the sensor was first at +40 mV to drive steady-state glucose oxidation, then the working electrode was electrically disconnected for a period of time (accumulation time) to accumulate charge, and the working electrode was reconnected and the accumulated charge was measured. As shown, the magnitude of the oxidation current spike increases with increasing accumulation time. Thus, simply increasing the accumulation time (e.g., to 30 seconds, 60 seconds, or 120 seconds) increases the sensitivity of the measurement with this glucose sensor and glucose concentration. In Figure 3D, the amperometric signal measured as the steady-state sensor current and the peak height and peak area of ​​the current spike measured in Figure 3C are plotted against accumulation time. As shown, the amperometric current is constant independent of accumulation time. However, both the height and area of ​​the current spikes show a linear dependence on the accumulation time, highlighting the advantage of accumulation mode sensing over conventional current measurements.That is, the sensitivity of the sensor may be adjusted by varying an easily adjustable parameter of the measurement technique (eg, the time for charge accumulation).

[0088] According to an embodiment of the present disclosure, the accumulation mode sensing method provides a signal over a range of analyte concentrations. Figures 4A and 4B show an example of a calibration experiment using an exemplary glucose sensor with glucose concentrations up to 100 μM. As shown, an accumulation time of 60 seconds was used for each detection. Figure 4A shows the resulting current waveform versus time for this experiment. As shown, the steady-state amperometric current and the magnitude of the accumulation mode current peak increase with increasing glucose concentration. Figure 4B shows a plot of the amperometric current as a function of glucose concentration and the peak height and peak area of ​​the current spike. All three signals show a linear dependence on analyte concentration. Thus, these results show that accumulation mode sensing measured using peak height or peak area can produce a linear calibration curve, and therefore accumulation mode sensing may be utilized for sensing in a manner similar to conventional amperometric measurements, with increased sensitivity. Thus, since the peak heights obtained from accumulation mode sensing are measured in current units, the sensitivity of this measurement method may be quantitatively compared to that of amperometric measurements. For example, the sensitivity of the measurement methods may be compared by comparing the slopes of the calibration curves shown in Figure 4B. By comparison, amperometric measurements have a sensitivity of 0.44nA / μM, whereas accumulation mode sensing (using peak height measurements) has a sensitivity of 1.69nA / μM, thus increasing the sensitivity of electrochemical measurements by approximately 4 times over amperometric measurements for an accumulation time of 60 seconds.

[0089] Furthermore, in some embodiments of the present disclosure, both the peak height and the peak area provide the same results and sensitivity. The means for measuring the resulting current signal at the working electrode includes calculating the peak height and / or the peak area.

[0090] In some embodiments of the present disclosure, sensors with an outer membrane are used for accumulation mode sensing. Because electrochemical sensors are often covered with an outer membrane (e.g., a polymeric membrane) to provide stability, mass transport limitation, biocompatibility to the sensing reagent, and / or to prevent fouling to the electrode, the polymer-coated sensor was tested to ensure that accumulation mode sensing was performed as expected. With reference to FIG. 5, a calibration curve was obtained by amperometry and accumulation mode sensing at glucose concentrations of 0 μM, 50 μM, 100 μM, 200 μM, and 500 μM using a glucose sensor coated with a flow-limiting polymeric outer membrane. Four consecutive measurements were performed at each glucose concentration with different accumulation times (1 min, 2 min, 5 min, and 10 min) as shown by the data points in FIG. 5.

[0091] As shown in FIG. 5, a linear response to analyte concentration was obtained from amperometric (left graph) and accumulation mode measurements (middle and right graphs). As expected, using amperometric (left graph in FIG. 5), the sensitivity of the sensor is independent of accumulation time. However, using accumulation mode sensing (middle and right graphs in FIG. 5), the sensitivity of the sensor increases with increasing accumulation time. Due to the flow-limiting outer membrane, the sensitivity of the sensor using both amperometric and accumulation mode sensing is much lower than the sensitivity of the sensor without an outer membrane. This is expected because the outer membrane limits the diffusion of the analyte to the sensing reagent. However, as shown in FIG. 5, accumulation mode sensing performed as expected when a polymer outer membrane was added to the sensor, providing another example of how the sensitivity of the sensor can be adjusted by changing the accumulation time. It should be noted that a predetermined period for charge accumulation using accumulation mode sensing with a continuous monitoring sensor exceeding 10 minutes can adversely affect the time resolution performance of the sensor. Thus, in some embodiments of the present disclosure, accumulation mode sensing is performed using a sensor having an outer membrane with a predetermined period for charge accumulation of up to 10 minutes.

[0092] Additionally, an outer membrane, such as a flow-limiting outer membrane, is not necessary to prevent fouling of the electrode when measuring analytes at low concentrations, but the outer membrane may act as a biocompatible interface with the in vivo environment and / or stabilize the underlying sensing layer (containing an electron transfer agent and / or analyte-specific enzyme thereon). In accumulation mode sensing where an outer membrane is used, the predetermined period of time for charge accumulation may be increased to oxidize the total concentration of the analyte. In some embodiments of the present disclosure, the accumulation mode sensing method using a sensor having an outer membrane includes increasing the predetermined period of charge accumulation to 1, 2, 3, 4, 5, 6, 7, 8, 9, or 10 minutes to fully react all analytes present at the working electrode. In some embodiments of the present disclosure, the accumulation mode sensing method using a sensor having an outer membrane includes increasing the predetermined period of charge accumulation to 10 to 30 minutes.

[0093] Alternatively, in some embodiments of the present disclosure, the outer membrane may be formed of a highly permeable material, which may allow for stability, limited mass transport, and / or biocompatibility without slowing down the rate at which the analyte reaches the sensing layer of the working electrode. Non-limiting examples of highly permeable membrane materials include polyvinylpyridine crosslinked with high molecular weight (MW: 400 g / mol or higher) polyethylene glycol diglycidyl ether, derivatized polyvinylpyridine crosslinked with high molecular weight (MW: 400 g / mol or higher) polyethylene glycol diglycidyl ether, polyvinyl alcohol, polyacrylic acid, and polymethacrylic acid.

[0094] With reference to FIGS. 6A-6B, an electrochemical glucose sensor was used in an in vitro experiment to measure (e.g., sense) glucose at concentrations ranging from 0 to 1000 nanomolar (nM). In this example, the working electrode of the sensor included a glucose oxidase enzyme crosslinked to an Os-based redox polymer. The Os-based redox polymer was deposited and immobilized on a screen-printed carbon electrode. The experiment was performed as disclosed herein (e.g., Example 8). A screen-printed carbon counter electrode and an Ag / AgCl reference electrode were also used. Prior to each measurement, the working electrode was held at +40 mV vs. Ag / AgCl for 3 minutes, after which the open circuit potential of the electrode was measured for 3 minutes. The graph in FIG. 6A shows the change in potential versus time obtained for the indicated glucose concentrations (0 to 1000 nM glucose). Thus, as shown, the higher the glucose concentration, the higher the potential drift rate. In some embodiments of the present disclosure, the drift rate is calculated as the slope of the potential versus time. Figure 6B is a calibration curve showing a plot of the drift rate (calculated as the slope from 30 to 180 s) versus glucose concentration. As shown in Figure 6B, the drift rate of potential shows a linear dependence on the glucose concentration.

[0095] 6C-6D, the same electrochemical glucose sensor used in the experiment of FIG. 6A-6B was used in an in vitro experiment to measure concentrations of glucose ranging from 0 to 750 nM, including concentrations less than 100 nM (e.g., 10 nM, 25 nM, and 50 nM). The graph in FIG. 6C shows the change in potential versus time obtained for the indicated glucose concentrations. Thus, as shown in FIG. 6D, the drift rate plotted in this experiment remains linear until the glucose concentration is reduced to 10 nM. This correlation is further illustrated in FIG. 6E, which shows the calibration curve obtained from testing eight individual glucose sensors. A control sensor without glucose oxidase enzyme (but with Os redox polymer) was also tested in this experiment. As shown in FIG. 6E and FIG. 6F, the drift rate of the control sensor, represented by open circles, did not show any dependence on glucose concentration.

[0096] In some embodiments of the present disclosure, the presently disclosed methods may be used to reduce background signals (e.g., signals at [analyte]=0). With reference to Figures 6G-6H, experiments were performed using the glucose sensor used in the experiment shown in Figure 6A as the working electrode. Also, a control sensor that does not contain glucose oxidase enzyme but does contain Os redox polymer was used as the reference electrode during open circuit potential measurements. This configuration is used to minimize the amount of signal measured that does not originate from glucose oxidation. For example, when a control sensor that does not contain glucose oxidase is used as the reference electrode, the background signal (slope of potential versus time when glucose concentration is zero) is nearly zero. The intercept of the calibration curve shown in Figure 6H is two orders of magnitude smaller than the intercept of the calibration curve shown in Figure 6F obtained using an Ag / AgCl reference electrode. Thus, the methods and systems of the present disclosure include using a control sensor that does not contain glucose oxidase as the reference electrode during open circuit potential measurements as an effective method of reducing background.

[0097] In some embodiments of the present disclosure, the signal generated from the redox reaction of the analyte at the sensing layer of the working electrode may be adjusted or modified to enhance the signal output for any given sensor and / or analyte concentration. In some embodiments of the present disclosure, the signal is enhanced by modifying the frequency at which the current signal is recorded. For example, as shown in FIG. 7, to maximize the peak height measured in the accumulation detection current spike, the signal may be recorded at a faster sampling rate (e.g., 0.1 Hz) and filtered at a higher frequency (e.g., 3.2 Hz) compared to the 0.5 Hz sampling rate and 0.03 Hz filter frequency used in the accumulation mode sensing experiments disclosed herein and shown in FIGS. 3A-3D, 4A-4B, and 5. As shown in FIG. 7, the detection peaks are much steeper and larger peak heights are obtained at frequencies higher than 3.2 Hz. Thus, in some embodiments of the present disclosure, the accumulation mode sensing method includes increasing the filter frequency to 3.2 Hz to maximize the signal magnitude. However, at frequencies higher than 3.2 Hz, the signal to noise ratio becomes too large to allow accurate measurements using amperometric current or accumulated peak measurements.

[0098] In some embodiments of the present disclosure, carbon nanotubes (CNTs) are added to the sensing element of the working electrode. For example, CNTs are added to a sensing reagent containing a redox mediator and an analyte-specific enzyme, and then applied to the working electrode. With reference to FIG. 8A, in the right micrograph, CNTs were added to the sensing reagent, whereas in the left micrograph, no CNTs were added. Accumulation mode sensing was measured with and without CNTs. As shown in FIG. 8B, the peak height of the accumulation mode current spike is larger when CNTs are added to the sensing element on the surface of the working electrode.

[0099] In some embodiments of the present disclosure, accumulation mode sensing includes using a sensor with an accumulation time (e.g., a predetermined period of time for charge accumulation), a signal frequency filter of 3.2 Hz, and a sensing element on the working electrode surface that includes carbon nanotubes (CNTs). FIG. 9A shows an accumulation mode signal obtained for a representative glucose sensor in the presence of CNTs, with glucose concentrations of 0-200 nM, with an accumulation time of 30 minutes, and with the signal filtered at 3.2 Hz. Thus, as shown in the signal calibration curve of FIG. 9B, accumulation mode sensing according to embodiments of the present disclosure provides enhanced sensitivity for low concentration analytes compared to amperometric measurements. As can be seen, for an accumulation time of 30 minutes, accumulation mode sensing using peak height measurements provides an 800-fold increase in sensitivity compared to amperometric measurements. In terms of detection limit, accumulation mode sensing using peak area measurements outperforms amperometric measurements, providing a lower limit of detection (LOD) of 4.7±1.4 nM, which is a 25-fold improvement over amperometric measurements. The linear range of sensing in accumulation mode is more limited than that of amperometry, but this range may be shifted to higher concentrations by using shorter accumulation times.

[0100] Accumulation mode detection sensor The sensors described herein may be in vivo sensors or in vitro sensors (i.e., separate monitoring strips). Such sensors may be formed on a substrate, e.g., a substantially planar substrate. In certain embodiments, the sensor is a wire, e.g., a working electrode wire inner portion having one or more other electrodes associated with (e.g., including wrapping around) the working electrode. The sensor may also include at least one counter electrode (or counter / reference electrode) and / or at least one reference electrode or at least one reference / counter electrode.

[0101] 12 shows a schematic of one embodiment of an analyte sensor 800 according to an embodiment of the present disclosure. The sensor includes electrodes 801, 802, and 803 on a substrate 804. The electrodes (and / or other structures) may be formed directly or via machining using any suitable technique (e.g., chemical vapor deposition (CVD), physical vapor deposition, sputtering, reactive sputtering, printing, coating, ablation (e.g., laser ablation), painting, dip coating, etching, etc.). Materials include, but are not limited to, any one or more of aluminum, carbon (including graphite), cobalt, copper, gallium, gold, indium, iridium, iron, lead, magnesium, mercury (as mercury alloys), nickel, niobium, osmium, palladium, platinum, rhenium, rhodium, selenium, silicon (e.g., doped polycrystalline silicon), silver, tantalum, tin, titanium, tungsten, uranium, vanadium, zinc, zirconium, mixtures thereof, and alloys, oxides, or metal compounds of these elements.

[0102] The analyte sensor 800 may be entirely implantable within a user, or may have only a portion located inside the user's body and another portion located outside the user's body. For example, the sensor 800 may include a first portion that may be positioned above the surface of the skin 810 and a second portion that may be positioned below the surface of the skin. In such an embodiment, the outer portion may include contacts (connected by wires to the electrodes of the second portion) and interface with other devices (e.g., a transmitter) outside the user. Although the embodiment of FIG. 12 shows three electrodes 801, 802, and 803 adjacently positioned on the same surface of the substrate 804, other configurations are contemplated. For example, more or fewer electrodes, some or all of the electrodes may be positioned on different surfaces of the substrate or on other substrates, some or all of the electrodes may be stacked, electrodes of different materials and sizes may be used, etc.

[0103] FIG. 13 illustrates a cross-sectional view of an embodiment of an analyte sensor 500 having a first portion (which may be characterized as a major portion in this embodiment) positionable above the surface of the skin and a second portion (which may be characterized as a minor portion in this embodiment) positionable below the surface of the skin (e.g., penetrating the skin (dermis) to the subcutaneous space and contacting the wearer's biological fluids (e.g., interstitial fluid)). The analyte sensor 500 includes a sensor tail 530 (also sometimes referred to herein as an insertion tip). Electrode contacts (not shown) are on the first portion of the sensor 500 that is positioned above the skin surface and extend to a location on the sensor tail 530. A working electrode 501, a reference electrode 502, and a counter electrode 503 are shown on the second portion of the sensor 500, and specifically on the bottom of the sensor tail 530. It will be appreciated that the sensor may include more or fewer electrodes without departing from the scope of the present disclosure. For example, the sensor may include more than one working electrode, and the separate counter and reference electrodes may be a single counter / reference electrode.

[0104] 13, the sensor 500 includes a substrate (or substrate layer) 504 and a first conductive layer 508 (e.g., carbon, gold, etc.). The first conductive layer 508 is in electrical communication with a sensing area 509, which together define a working electrode 501. By imparting antimicrobial qualities to one or more components of the sensor 500, the sensing area 509 may be protected from microorganisms to protect the health of the wearer's skin and / or to protect the sensing area 509 from potential interference by such microorganisms (e.g., formation of a biofilm due to potential transfer of microorganisms). The various electrodes and sensing area defined at the bottom of the sensor tail 530 in FIG. 13 may collectively be referred to as the sensing area, and any antimicrobial qualities imparted to the sensor tail as described herein may be imparted to the top (top 25%) of the sensor tail 530 over said area (e.g., over the sensing area 509, or over the electrode 503).

[0105] In some embodiments, a first insulating layer 505, such as a first dielectric layer, may be provided or laminated on at least a portion of the first conductive layer 508, and a second conductive layer 511 may be provided or laminated on at least a portion of the first insulating (or dielectric) layer 505. The reference electrode 502 may be provided by the second conductive layer 511 and a second conductive material 510 (e.g., a layer of silver / silver chloride (Ag / AgCl)) as shown in Figure 13. Figure 14B shows another possible arrangement of the second conductive material 510, with an outer membrane 520 covering these layers.

[0106] In some embodiments, a second insulating layer 506, such as a second dielectric layer, may be applied or laminated to at least a portion of the second conductive layer 511. Additionally, a third conductive layer 513 may be applied to at least a portion of the second insulating layer 506 to provide a counter electrode 503. Finally, a third insulating layer 507 may be applied or laminated to at least a portion of the third conductive layer 513. In this manner, the sensor 500 may be laminated such that at least a portion of each of the conductive layers is separated by an insulating layer (e.g., a dielectric layer). Other layer configuration possibilities are shown in FIG. 14B. Although layers of different lengths are shown in the embodiments of FIG. 13 and FIG. 14B, some or all of these layers may have the same or different lengths and widths without departing from the scope of the present disclosure.

[0107] In any one or all of the embodiments, some or all of the electrodes 501, 502, and 503 may be provided on the same side of the substrate 504 of the laminated construction, or two or more electrodes may be provided coplanarly, i.e., adjacent, parallel, angled to each other, disposed on the same surface of the substrate 504. For example, coplanar electrodes may include suitable spacing therebetween and / or may include dielectric or insulating material provided between the conductive layers / electrodes. Furthermore, in some embodiments, one or more of the electrodes 501, 502, and 503 may be provided on different sides of the substrate 504. In such embodiments, the contact pads may be on the same or different sides of the substrate. For example, the electrodes may be on a first side and their contacts on a second side, e.g., the traces and contacts connecting the electrodes may cross the substrate.

[0108] 14A, another embodiment of an analyte sensor is shown that represents a variation of the sensor 500 of FIG. 13 and FIG. 14B in accordance with one or more embodiments of the present disclosure. Referring to FIG. 14A, an implantable (e.g., subcutaneous or transcutaneous) sensing region 920 is shown in accordance with one or more embodiments of the present disclosure that includes a working electrode 922 that includes a sensing element 931. A proximal end 940 is adapted to connect with various electrical connections for transmitting an output signal of the sensing region 920. The distal end 925 and the proximal end 940 together form a sensor tail. The sensing region 920 encompasses the bottom of the sensor tail. As shown, the sensing region 920 includes a rounded tip, but may have other tip shapes that allow for easy insertion into the wearer's skin.

[0109] Also, in one or more embodiments, the sensing area 920 may include a reference electrode, a counter electrode, or a counter / reference electrode, as shown in Figures 13 and 14B. Alternative electrode configurations may be used without departing from the scope of the present disclosure.

[0110] 13, 14A, and 14B, sensors (or sensing regions) 500, 920 include sensing functionality at a distal portion of their respective sensor tails. As discussed above, this location may provide increased contact deeper beneath the wearer's skin (e.g., the subcutaneous space) and closer access to the wearer's interstitial fluid and therefore the analyte (e.g., its concentration) being measured. That is, placing a sensing region deep enough within the wearer's skin may allow for accurate measurement of a particular analyte, while locating a sensing region closer to the skin surface may be inadequate to properly determine the concentration or other characteristic of the desired analyte.

[0111] 13 and 14B-14D, one or more embodiments of the present disclosure include a working electrode 501 or 320 that includes a sensing region 509. The sensing region 509 includes at least one sensing element 322 that includes, for example, an analyte-specific enzyme 323 and an electron transfer agent (e.g., a redox mediator) 324. The working electrode 501 or 320 is disposed on a substrate 504 or 325. The substrate 504 or 325 is disposed between and in contact with the working electrode 501 or 320 and a counter electrode 503. A first insulating layer 505 is disposed in contact with a surface of the working electrode 501 or 320 that is not in contact with the substrate 504 or 325. A reference electrode 502 is provided in contact with a surface of the first insulating layer 505 that is not in contact with the working electrode 501 or 320, and a second conductive material (or layer) 510 is provided in contact with a surface of the reference electrode 502 that is not in contact with the first insulating layer 505.

[0112] 14C, a sensing element 322 is provided on at least a portion of the working electrode 320. In some embodiments of the present disclosure, two or more sensing elements 322 may be provided on a sensing layer of the working electrode, the two or more sensing elements being provided laterally relative to one another.

[0113] In some embodiments of the present disclosure, any suitable configuration of sensing element 322 may be provided on the working electrode 320. Further configurations of sensing elements are disclosed, for example, in Hoss et al. (US 2012 / 0150005), the entire contents of which are incorporated herein by reference.

[0114] In some embodiments of the present disclosure, referring to FIG. 14B, the sensor 500 includes an outer membrane 520 that covers at least the working electrode 501 and the sensing area 509. In other embodiments, the outer membrane 520 covers the entire sensor 500. In some embodiments, the outer membrane 520 covers all of the active area of ​​the sensor 500. For example, the active area of ​​the sensor 500 is shown as sensing area 920 in FIG. 14A and sensing area 509 in FIG. 14B. In some embodiments, the outer membrane 520 covers the working, counter, and / or reference electrodes over the sensing area 920 or sensing area 509.

[0115] 14C shows an enlarged perspective view of the outer membrane 335. The outer membrane 335 covers the sensing element 322 disposed on the working electrode 320 disposed on the substrate 325. As shown, the outer membrane 335 is in the process of being covered. The outer membrane 335 covers at least the entire sensing element 322.

[0116] Analyte-specific enzymes and electron transfer agents (redox mediators) In some embodiments of the present disclosure, the sensor of the present disclosure cannot directly measure the analyte. That is, the electrodes on the sensor cannot directly interact with the analyte. Thus, the analyte is detected by an enzyme protein that can directly interact with the analyte molecule. However, some enzymes (e.g., glucose oxidase) cannot directly exchange electrons with the electrode because their redox active sites are buried deep within the enzyme protein structure. Thus, an electron transfer agent (i.e., a redox mediator) is used to transfer electrons between the redox active site of the enzyme and the electrode. Since the immobilized molecules can relay electrons, the electron transfer agent and the analyte-specific enzyme are immobilized on the sensing layer to form what is called "wiring." Thus, they are "electrically connected." The analyte-specific enzyme is also called a "connected enzyme." Linked enzymes are disclosed, for example, in Gregg et al. (U.S. Pat. No. 5,262,035), Say et al. (U.S. Pat. No. 6,134,461), and Hoss et al. (U.S. Patent Publication No. 2012 / 0150005), the entire contents of which are incorporated herein by reference. In some embodiments, the analyte-specific enzyme is crosslinked to an electron transfer agent.

[0117] In some embodiments of the present disclosure, the electron transfer agent (e.g., redox mediator) is an easily electroreducible and electrooxidizable ion or molecule with a redox potential (voltage) several hundred millivolts higher or lower than the redox potential (voltage) of a standard calorific electrode (SCE). In some embodiments, the electron transfer agent is reducible at about -150 mV and oxidizable at about +400 mV vs. SCE. Examples of suitable redox mediators in the form of redox polymers are disclosed, for example, in Mao et al. (U.S. Pat. No. 6,605,200), the entire contents of which are incorporated herein by reference.

[0118] According to an embodiment of the present disclosure, as shown in Figure 14D, an electron transfer agent 324 is immobilized on the working electrode 320. In some embodiments, both the electron transfer agent 324 and the analyte-specific enzyme 323 are immobilized on the working electrode 320 by any suitable means. In some embodiments, the electron transfer agent and the analyte-specific enzyme are co-immobilized on the working electrode by any suitable cross-linking agent. In some embodiments, the electron transfer agent and the analyte-specific enzyme are co-immobilized by a chemical cross-linking agent, such as, for example, polyethylene glycol diglycidyl ether (PEGDGE).

[0119] In some embodiments of the present disclosure, the electron transfer agent used for accumulation mode sensing comprises a redox species selected from osmium, ruthenium, iron, or cobalt bound to a polymer selected from polyvinylpyridine, polythiophene, polyaniline, polypyrrole, or polyacetylene. In some embodiments, the electron transfer agent is an osmium (Os)-containing polyvinylpyridine redox polymer of formula I: [ka]

[0120] In some embodiments of the present disclosure, the electron transfer agent may be an organic redox species, an organometallic redox species, or an inorganic redox species. Examples of organic redox species are quinones and species with quinoid structures in their oxidized state, such as Nile Blue and indophenol. Some quinones and partially oxidized quinhydrones react with functional groups of proteins, such as the thiol groups of cysteine, the amine groups of lysine and arginine, and the phenolic groups of tyrosine. These groups may make the redox species unsuitable for some of the sensors of the present disclosure due to the presence of interfering proteins in the analyte-containing fluid. However, most substituted quinones and molecules with quinoid structures have low reactivity with proteins. In some embodiments, tetra-substituted quinones have carbon atoms at the 1, 2, 3, and 4 positions.

[0121] Suitable electron transfer agents for use in the accumulation mode sensing method according to embodiments of the present disclosure have a structure or charge that prevents or substantially reduces diffusional loss of the electron transfer agent during analysis of a sample. In some embodiments of the present disclosure, the electron transfer agent comprises a polymer-bound redox species that can be immobilized on the sensing layer of the working electrode. The bond between the redox species and the polymer can be a covalent bond, a coordinate bond, or an ionic bond. Useful electron transfer agents and methods for making them are described in U.S. Patent Nos. 5,264,104; 5,356,786; 5,262,035; and 5,320,725, the entire contents of which are incorporated herein by reference. While either organic or organometallic redox species may be polymer-bound and used as the electron transfer agent, in some embodiments of the present disclosure, the redox mediator is a transition metal compound or complex. In some embodiments, the transition metal compound or complex includes osmium, ruthenium, iron, and cobalt compounds or complexes. Obviously, many of the redox mediator species described herein may be used as electron transfer agents, for example, without a polymeric component, provided that leaching of the electron transfer agent is acceptable in the carrier fluid or in the sensing layer of the sensor.

[0122] One class of non-emissive polymeric electron transfer agents contains redox species covalently bound to the polymer composition. One example of this type of mediator is polyvinylferrocene.

[0123] Another type of non-releasable polymeric electron transfer agent includes ionically bound redox species. Typically, this type of mediator includes a charged polymer bound to an oppositely charged redox species. Examples of this type of mediator include a negatively charged polymer (e.g., Nafion (Dupont)) bound to a positively charged redox species (e.g., polypyridyl cation bound to osmium, ruthenium, iron, or cobalt). Another example of an ionically bound mediator is a positively charged polymer (e.g., quaternized poly(4-vinylpyridine) or poly(1-vinylimidazole)) bound to a negatively charged redox species (e.g., ferricyanide or ferrocyanide). In some embodiments of the present disclosure, the bound redox species is a highly charged redox species bound within an oppositely charged redox polymer.

[0124] In another embodiment of the present disclosure, suitable non-emissive polymeric electron transfer agents include redox species coordinated to a polymer. For example, the mediator may be formed by coordinating 2,2'-dipyridyl complexes of osmium or cobalt to poly(1-vinylimidazole) or poly(4-vinylpyridine).

[0125] In some embodiments of the present disclosure, the electron transfer agent is an osmium transition metal complex having one or more ligands, each of which has a nitrogen-containing heterocycle (e.g., 2,2'-dipyridine, 1,10-phenanthroline, or derivatives thereof). Additionally, in some embodiments, the electron transfer agent has one or more ligands covalently attached to a polymer, each of which has at least one nitrogen-containing heterocycle (e.g., pyridine, imidazole, or derivatives thereof). These preferred electron transfer agents rapidly exchange electrons between the respective electron transfer agents and the working electrode, allowing rapid oxidation or reduction of the complex.

[0126] In some embodiments of the present disclosure, the electron transfer agent comprises (a) a polymer or copolymer having pyridine or imidazole functional groups, and (b) an osmium cation complexed with two ligands, each of which comprises 2,2'-dipyridine, 1,10-phenanthroline, or a derivative thereof, and the two ligands are not necessarily the same. In some embodiments, the derivatives of 2,2'-dipyridine that are complexed with the osmium cation include 4,4'-dimethyl-2,2'-dipyridine, and mono-, di-, and polyalkoxy-2,2'-dipyridines (e.g., 4,4'-dimethoxy-2,2'-dipyridine). In some embodiments, the derivatives of 1,10-phenanthroline that are complexed with the osmium cation are 4,7-dimethyl-1,10-phenanthroline, and mono-, di-, and polyalkoxy-1,10-phenanthrolines (e.g., 4,7-dimethoxy-1,10-phenanthroline). In some embodiments of the present disclosure, the polymers that are complexed with the osmium cation are polymers and copolymers of poly(1-vinylimidazole) (referred to as "PVI") and poly(4-vinylpyridine) (referred to as "PVP"). Suitable copolymer substituents of poly(1-vinylimidazole) include acrylonitrile, acrylamide, and substituted or quaternized N-vinylimidazole. In some embodiments, the electron transfer agent comprises osmium complexed with a polymer or copolymer of poly(1-vinylimidazole).

[0127] According to embodiments of the present disclosure, the electron transfer agent has a redox potential (voltage) in the range of -100 mV to about +150 mV vs. the standard calorific electrode (SCE). More specifically, the potential (voltage) of the electron transfer agent is in the range of -100 mV to +150 mV. In some embodiments, the potential (voltage) is in the range of -50 mV to +50 mV. In other embodiments of the present disclosure, the electron transfer agent has osmium, ruthenium, iron, or cobalt redox centers and has a redox potential (voltage) in the range of +50 mV to -150 mV vs. the SCE.

[0128] Examples of analyte-specific enzymes In some embodiments of the present disclosure, an analyte-specific enzyme is provided (e.g., immobilized) on the working electrode surface to catalyze the oxidation of the analyte to be measured. As used herein, the analyte-specific enzyme may also be referred to as an analyte oxidase. In some embodiments of the present disclosure, the analyte-specific enzyme is selected from glucose oxidase, NAD-glucose dehydrogenase, and FAD-glucose dehydrogenase, which oxidizes glucose. In some embodiments, the analyte-specific enzyme is lactate oxidase or NAD-lactate dehydrogenase, which oxidizes lactate. In some embodiments, the analyte-specific enzyme is NAD-3-hydroxybutyrate dehydrogenase, which oxidizes 3-hydroxybutyrate. In some embodiments, the analyte-specific enzyme is 11β-hydroxysteroid dehydrogenase type 2, which oxidizes cortisol. In some embodiments, the analyte-specific enzyme is NAD-alcohol dehydrogenase, which oxidizes alcohol. In some embodiments, the analyte-specific enzyme is pyruvate oxidase, which oxidizes pyruvate. In some embodiments, the analyte specific enzyme is NAD-glutamate dehydrogenase, which oxidizes glutamate.In some embodiments, the analyte specific enzyme is xanthine oxidase, which oxidizes theophylline.

[0129] As will be apparent to one of skill in the art, any nicotinamide adenine dinucleotide (NAD) or flavin oxidase enzyme may be bound or immobilized on the sensing layer of the working electrode to oxidize a substrate of a corresponding analyte.

[0130] In some embodiments of the disclosure, examples of NAD-dependent enzymes include (-)-borneol dehydrogenase, (+)-borneol dehydrogenase, (+)-sabinol dehydrogenase, (+)-trans-carveol dehydrogenase, (3S,4R)-3,4-dihydroxycyclohexa-1,5-diene-1,4-dicarboxylic acid dehydrogenase, (R,R)-butanediol dehydrogenase, (R)-2-hydroxyfatty acid dehydrogenase, (R)-2-hydroxyacid dehydrogenase, (R)-4-hydroxyphenyl triphosphate ... hydroxyphenyllactate dehydrogenase, (R)-aminopropanol dehydrogenase, (R)-dehydropantoate dehydrogenase, (S,S)-butanediol dehydrogenase, (S)-2-hydroxyfatty acid dehydrogenase, (S)-carnitine-3-dehydrogenase, (S)-usnic acid reductase, 1,2-dihydroxy-6-methylcyclohexa-3,5-diene carboxylate dehydrogenase, 1,3-propanediol dehydrogenase, 1,6-dihydroxycyclohexa-2,4-diene-1-carboxylate dehydrogenase genase, 2-(R)-hydroxypropyl-CoM dehydrogenase, 2-(S)-hydroxypropyl-CoM dehydrogenase, 2-alkenal reductase, 2-alkyne-1-ol dehydrogenase, 2-aminobenzenesulfonate-2,3-dioxygenase, 2-chlorobenzoate-1,2-dioxygenase, 2-coumarate reductase, 2-dehydro-3-deoxy-D-gluconate-5-dehydrogenase, 2-deoxy-D-gluconate-3-dehydrogenase, 2-enoate reductase, 2-hydroxy-1,4-benzoate reductase Quinone reductase, 2-hydroxy-3-oxopropionic acid reductase, 2-hydroxybiphenyl-3-monooxygenase, 2-hydroxymethylglutarate dehydrogenase, 2-hydroxyquinoline-5,6-dioxygenase, 2-hydroxyquinoline-8-monooxygenase, 2-oxoadipate reductase, 2-oxoaldehyde dehydrogenase (NAD+), 2-oxoisovalerate dehydrogenase (acylation), 2,3-dihydro-2,3-dihydroxybenzoate dehydrogenase, 2,3-dihydroxy-2,3-Dihydro-p-cumic acid dehydrogenase, 2,4-diaminopentanoic acid dehydrogenase, 2,6-dihydroxypyridine-3-monooxygenase, 2'-phosphotransferase, 3-(imidazol-5-yl)lactate dehydrogenase, 3"-deamino-3"-oxonicotianamine reductase, 3-dehydro-L-gulonic acid 2-dehydrogenase, 3-hydroxy-2-methylbutyryl-CoA dehydrogenase, 3-hydroxy-2-methylpyridine carboxylate dioxygenase, 3-hydroxyacyl-CoA dehydrogenase, 3-hydroxybenzoic acid 6-monooxygenase, 3-hydroxybutyric acid dehydrogenase, 3-hydroxyisobutyric acid dehydrogenase, 3-hydroxyphenylacetate 6-hydroxylase, 3-hydroxypimeloyl-CoA dehydrogenase, 3-hydroxypropionic acid dehydrogenase, 3-methylbutanal reductase ductase, 3-oxoacyl-(acyl carrier protein) reductase (NADH), 3-phenylpropanoate dioxygenase, 3(or 17)a-hydroxysteroid dehydrogenase, 3α-hydroxy-5β-androstan-17-one-3α-dehydrogenase, 3α-hydroxycholanic acid dehydrogenase, 3α-hydroxysteroid dehydrogenase (A-specific), 3α-hydroxysteroid dehydrogenase (B-specific), 3α,7α,12α-trihydroxycholestane-26-al-26-oxidoreductase, 3α(17β)-hydroxysteroid dehydrogenase (NAD+), 3α(or 20β)-hydroxysteroid dehydrogenase, 3β-hydroxysteroid dehydrogenase, 4-(hydroxymethyl)benzenesulfonic acid dehydrogenase, 4-aminobenzoate-1-monooxygenase, 4-chlorophenylacetate-3,4-Dioxygenase, 4-formylbenzenesulfonate dehydrogenase, 4-hydroxy-tetrahydrodipicolinate reductase, 4-hydroxybenzaldehyde dehydrogenase, 4-hydroxybenzoate-1-hydroxylase, 4-hydroxybenzoate-3-monooxygenase (NAD(P)H), 4-hydroxybutyrate dehydrogenase, 4-hydroxycyclohexanecarboxylate dehydrogenase, 4-hydroxymuconic acid-semialdehyde dehydrogenase, 4-hydroxyphenylacetaldehyde dehydrogenase 4-Hydroxyphenylacetate-1-monooxygenase, 4-Hydroxyquinoline-3-monooxygenase, 4-Hydroxythreonine-4-phosphate dehydrogenase, 4-Nitrophenol-2-monooxygenase, 4-Oxoproline reductase, 4-Phosphoerythronic acid dehydrogenase, 4-Sulfobenzoate-3,4-dioxygenase, 4-Trimethylammoniobutyraldehyde dehydrogenase, 5-Carboxymethyl-2-hydroxymuconic acid-semialdehyde dehydrogenase, 5,6-Dihydroxy-3- Methyl-2-oxo-1,2,5,6-tetrahydroquinoline dehydrogenase, 6-endo-hydroxycineole dehydrogenase, 6-hydroxyhexanoic acid dehydrogenase, 6,7-dihydropteridine reductase, 7-α-hydroxysteroid dehydrogenase, 15-hydroxyicosatetraenoic acid dehydrogenase, 15-hydroxyprostaglandin dehydrogenase (NAD+), 15-oxoprostaglandin-13-oxidase, 16-α-hydroxysteroid dehydrogenase, 17β-hydroxysteroid dehydrogenase, Cytosteroid dehydrogenase, 20-α-hydroxysteroid dehydrogenase, 21-hydroxysteroid dehydrogenase (NAD+), ADP-glyceromanno-heptose-6-epimerase, alanine dehydrogenase, alanopine dehydrogenase, alcohol dehydrogenase, alcohol dehydrogenase (NAD(P)+), aldehyde dehydrogenase (NAD(P)+), aldehyde dehydrogenase (NAD+), aldose-1-dehydrogenase, alkene monooxygenase, α-santonin-1,2-reductase, aminobutyraldehyde dehydrogenase, aminomuconate-semialdehyde dehydrogenase, anthocyanidin reductase, anthranilate-1,2-dioxygenase (deaminating, decarboxylating), anthraniloyl-CoA monooxygenase, apiose-1-reductase, aquacobalamin reductase, arogenate dehydrogenase, arogenate dehydrogenase (NAD(P)+), aryl-alcohol dehydrogenase, Aryl-aldehyde dehydrogenase, asparagusate reductase, aspartate dehydrogenase, ATP-dependent NAD(P)H-hydrate dehydratase, benzaldehyde dehydrogenase (NAD+), benzene-1,2-dioxygenase, benzoate-1,2-dioxygenase, β-alanopine dehydrogenase, betaine-aldehyde dehydrogenase, biphenyl-2,3-dioxygenase, butanal dehydrogenase, carnitine-3-dehydro genase, CDP-4-dehydro-6-deoxyglucose reductase, CDP-glucose-4,6-dehydratase, CDP-paratose-2-epimerase, cholest-5-ene-3β,7α-diol-3β-dehydrogenase, cholestanetetraol-26-dehydrogenase, cis-1,2-dihydro-1,2-dihydroxynaphthalene dehydrogenase, cis-1,2-dihydrobenzene-1,2-diol dehydrogenase, cis-1,2-dihydroxamyl dehydrogenase, cis-4-methylcyclohexa-3,5-diene-1-carboxylate dehydrogenase, cis-2,3-dihydrobiphenyl-2,3-diol dehydrogenase, cis-3,4-dihydrophenanthrene-3,4-diol dehydrogenase, cis-dihydroethylcatechol dehydrogenase, CoA-disulfide reductase, cobalamin(II) reductase, coniferyl-aldehyde dehydrogenase, cucurbitacin δ23-reductase, cyclohexane-1,2-diol dehydrogenase, cyclohexanol dehydrogenase, cyclopentanol dehydrogenase, cystine reductase, D-arabinitol-2-dehydrogenase, D-arabinitol-4-dehydrogenase, D-arabinose-1-dehydrogenase, D-arabinose-1-dehydrogenase (NAD(P)+), D-iditol-2-dehydrogenase, D-malate dehydrogenase (decarboxylating), D-threo-aldose-1-dehydrogenase, D-xylose-1-dehydrogenase, D-xylulose reductase, dibenzothiophene dihydrodiol dehydrogenase. Diferric transferrin reductase, dihydrouracil dehydrogenase (NAD+), diiodophenylpyruvate reductase, dimethylmalate dehydrogenase, DTDP-glucose-4,6-dehydratase, ephedrine dehydrogenase, erythrose-4-phosphate dehydrogenase, estradiol-17α-dehydrogenase, estradiol-17β-dehydrogenase, fatty acid-acyl-CoA synthase, ferredoxin-NAD(+) reductase, ferric chelate reductase, fluoren-9-ol dehydrogenase, fluoroacetaldehyde dehydrogenase hydrogenase, FMN reductase, formaldehyde dehydrogenase, fructuronic acid reductase, fumarate reductase (NADH), furylfuramide isomerase, galactitol-2-dehydrogenase, galactitol-1-phosphate-5-dehydrogenase, galactose-1-dehydrogenase, gamma-guanidinobutyraldehyde dehydrogenase, GDP-4-dehydro-6-deoxy-D-mannose reductase, GDP-4-dehydro-D-rhamnose reductase, GDP-6-deoxy-D-talose-4-dehydrogenase, GDP-mannose-4,6-dehydratase, GDP-mannose-6-dehydrogenase, gluconate-5-dehydrogenase, glucose-1-dehydrogenase, glucose-1-dehydrogenase (NAD+), glutamate synthase (NADH), glutarate-semialdehyde dehydrogenase, glyceraldehyde-3-phosphate dehydrogenase (NAD(P)+), glyceraldehyde-3-phosphate dehydrogenase (phosphorylation), glycerate dehydrogenase, glycerol dehydrogenase, glycerol-3-phosphate dehydrogenase (NAD(P)+), glycerol-3-phosphate dehydrogenase (NAD+), glycine cleavage system, glycine dehydrogenase, glycolaldehyde dehydrogenase, glyoxylate reductase, hexadecanal dehydrogenase (acylation) , hexadecanol dehydrogenase, histidinol dehydrogenase, homoisocitrate dehydrogenase, homoserine dehydrogenase, hydrogen dehydrogenase, hydroxycyclohexanecarboxylate dehydrogenase, hydroxylamine reductase (NADH), hydroxymalonate dehydrogenase, hydroxymethylglutaryl-CoA reductase, hydroxyphenylpyruvate reductase, hydroxypyruvate reductase, hyponitrite reductase, hypotaurine dehydrogenase, icosanoyl-CoA synthase, imidazoleacetate-4-monooxygenase, IMP dehydrogenase, indanol dehydrogenase, indole-3-acetaldehyde reductase (NADH), indolelactate dehydrogenase, inositol-2-dehydrogenase, , inositol-3-phosphate synthase, isocitrate dehydrogenase, isopiperitenol dehydrogenase, kynurenic acid-7,8-dihydrodiol dehydrogenase, L-amino acid dehydrogenase, L-aminoadipate-semialdehyde dehydrogenase, L-arabinitol-2-dehydrogenase, L-arabinitol-4-dehydrogenase, L-arabinose-1-dehydrogenase, L-erythro-3,5-diaminohexanoate dehydrogenase, L-glycol dehydrogenase, L-gulonic acid-3-dehydrogenase, L- Iditol-2-dehydrogenase, L-idonic acid-5-dehydrogenase, L-rhamnose-1-dehydrogenase, L-threonic acid-3-dehydrogenase, L-threonine-3-dehydrogenase, lactaldehyde dehydrogenase, lactaldehyde reductase, lathosterol oxidase, leghemoglobin reductase, leucine dehydrogenase, long-chain alcohol dehydrogenase, lysine dehydrogenase, malate dehydrogenase (decarboxylation), malate dehydrogenase (oxaloacetate-decarboxylation), malate dehydrogenase (decarboxylation) Leuylacetate reductase, malonate-semialdehyde dehydrogenase, malonate-semialdehyde dehydrogenase (acetylation), mannitol-2-dehydrogenase, mannitol dehydrogenase, mannitol-1-phosphate-5-dehydrogenase, mannuronate reductase, melilotate-3-monooxygenase, mesotartarate dehydrogenase, methanol dehydrogenase, methylenetetrahydrofolate dehydrogenase (NAD+), methylglyoxal reductase (NADH-dependent), methylmalonate-semialdehyde dehydrogenase dehydrogenase (acylation), mevaldate reductase, monodehydroascorbate reductase (NADH), morphine-6-dehydrogenase, mycothiol-dependent formaldehyde dehydrogenase, mycothione reductase, myristoyl-CoA-11-(E) desaturase, myristoyl-CoA-11-(Z) desaturase, N-acetylhexosamine-1-dehydrogenase, N-acylmannosamine-1-dehydrogenase, N-hydroxy-2-acetamidofluorene reductase, NAD(+)-dinitrogen-reductase,ADP-D-ribosyltransferase, NAD(+)-diphthamide-ADP-ribosyltransferase, NAD(P)(+)-protein-arginine-ADP-ribosyltransferase, NAD(P)+ nucleosidase, NAD(P)+ transhydrogenase (Re / Si-specific), NAD(P)+ transhydrogenase (Si-specific), NAD(P)H dehydrogenase (quinone 1), NAD(P)H dehydrogenase (quinone), NAD+ diphosphatase, NAD+ nucleosidase, NAD+ synthase, NA D+ synthase (glutamine hydrolysis), NADH dehydrogenase (quinone), NADH peroxidase, naphthalene-1,2-dioxygenase, nicotinamide-nucleotide adenylyltransferase, nitric oxide dioxygenase, nitrite reductase (NAD(P)H), nitroquinoline-N-oxide reductase, octanol dehydrogenase, ω-hydroxydecanoate dehydrogenase, opine dehydrogenase, orcinol-2-monooxygenase, ornithine cyclodeaminase, orotate reductase (NA DH), oxaloglycolate reductase (decarboxylation), pantoate-4-dehydrogenase, perillyl alcohol dehydrogenase, phenylacetaldehyde dehydrogenase, phenylalanine dehydrogenase, phenylglyoxylate dehydrogenase (acylation), phosphatidylcholine-12-monooxygenase, phosphatidylcholine desaturase, phosphogluconate-2-dehydrogenase, phosphoglycerate dehydrogenase, phosphonate dehydrogenase, phthalate-4,5-cis-dihydrodiol dehydrogenase hydrogenase, phthalate-4,5-dioxygenase, pimeloyl-CoA dehydrogenase, precorrin-2 hydrogenase, precorrin-3B synthase, prephenate dehydrogenase, propanediol-phosphate dehydrogenase, protein-disulfide reductase, pyridoxal-4-dehydrogenase, pyrroline-2-carboxylate reductase, pyrroline-5-carboxylate reductase, quinate dehydrogenase, retinal dehydrogenase, retinol dehydrogenase, ribitol-2-dehydrogenase,Ribitol-5-phosphate-2-dehydrogenase, rubredoxin-NAD(+) reductase, rubredoxin-NAD(P)(+) reductase, S-(hydroxymethyl)glutathione dehydrogenase, saccharopine dehydrogenase (NAD+, L-glutamate forming), saccharopine dehydrogenase (NAD+, L-lysine forming), salicylaldehyde dehydrogenase, salicylate-1-monooxygenase, sequoyitol dehydrogenase, serine-2-dehydrogenase, Sn-glycerol -1-phosphate dehydrogenase, sorbitol-6-phosphate-2-dehydrogenase, steroid-17α-monooxygenase, sterol-4α-carboxylate-3-dehydrogenase (decarboxylation), strombine dehydrogenase, succinic semialdehyde dehydrogenase, succinic semialdehyde dehydrogenase (NAD(P)+), succinylglutamic semialdehyde dehydrogenase, sulcatone reductase, tagaturonic acid reductase, tartrate dehydrogenase, tauropine Dehydrogenase, Taxifolin-8-monooxygenase, Terephthalate-1,2-cis-dihydrodiol dehydrogenase, Terephthalate-1,2-dioxygenase, Testosterone-17β-dehydrogenase, Tetrahydroxypteridine cycloisomerase, Thiomorpholine-carboxylate dehydrogenase, TM0436, Toluene dioxygenase, Trans-2-enoyl-CoA reductase (NAD+), Trimethylamine-N-oxide reductase, Tryptophan dehydrogenase, U DP-glucose-4-epimerase, UDP-glucose-6-dehydrogenase, UDP-glucuronic acid-5'-epimerase, UDP-glucuronic acid decarboxylase, UDP-N-acetylglucosamine-6-dehydrogenase, ureidoglycolate dehydrogenase, uronate dehydrogenase, vanillic acid monooxygenase, vanillin dehydrogenase, vomifoliol dehydrogenase, xanthine dehydrogenase, xanthomatine reductase, or xanthoxin dehydrogenase.

[0131] In some embodiments of the present disclosure, the analyte specific enzyme comprises a flavin oxidase, such as a flavin adenine dinucleotide (FAD) dependent oxidase or a flavin mononucleotide (FMN) dependent oxidase. Examples of FAD-dependent or FMN-dependent oxidases include (R)-6-hydroxynicotine oxidase, (S)-2-hydroxyacid oxidase, (S)-6-hydroxynicotine oxidase, 2-enoic acid reductase, 2-methyl-branched-enoyl-CoA reductase, 2-nitropropane dioxygenase, 2,4-dichlorophenol-6-monooxygenase, 2,6-dihydroxypyridine-3-monooxygenase, 3-aci-nitropropanoate oxidase, 3-hydroxy-2-methylpyridine carboxylate dioxygenase, 3-hydroxybenzoate-4-monooxygenase, 3-hydroxybenzoate-6-monooxygenase, 3-hydroxyphenylacetate-6-hydroxylase, 4-aminobenzoate-1-monooxygenase, 4-cresol dehydrogenase (hydroxy 4-Hydroxybenzoate-1-hydroxylase, 4-Hydroxybenzoate-3-monooxygenase, 4-Hydroxybenzoate-3-monooxygenase (NAD(P)H), 4-Hydroxymandelate oxidase, 4-Hydroxyphenylacetate-1-monooxygenase, 4-Hydroxyphenylacetate-3-monooxygenase, 4-Nitrophenol-2-monooxygenase, 4-Sulfobenzoate-3,4-dioxygenase, 5-Pyridoxic acid dioxygenase, Acyl-CoA oxidase, Adenylyl-sulfate reductase, Albendazole monooxygenase, Alcohol oxidase, Anthraniloyl-CoA monooxygenase, Aquacobalamin reductase, Aquacobalamin reductase (NADPH), Arginine-2-monooxygenase, Benzene-1,2-dioxygenase, Benzoate-1,2-dioxygenase, β-cyclopiazonic acid dehydrogenase, cellobiose dehydrogenase (receptor), choline oxidase, CoA-glutathione reductase, cobalamin(II) reductase, cyanocobalamin reductase (cyanide removal), cyclohexylamine oxidase, D-2-hydroxyacid dehydrogenase, D-amino acid oxidase, D-arabinono-1,4-lactone oxidase, D-aspartate oxidase, D-glutamate (D-aspartate) oxidase, D-lactate dehydrogenase (cytochrome), D-sorbitol dehydrogenase (receptor), dehydrogluconate dehydrogenase, deoxyribodipyri mide photolyase, dihydrouracil oxidase, dimethylamine dehydrogenase, dimethylglycine dehydrogenase, dimethylglycine oxidase, ferredoxin-NADP(+) reductase, gluconate-2-dehydrogenase (acceptor), glucose dehydrogenase (acceptor), glucoside-3-dehydrogenase, glutamate synthase (ferredoxin), glutamate synthase (NADH), glutamate synthase (NADPH), glutathione oxidase, glycerol-3-phosphate oxidase, hydrogen dehydrogenase, hydroxylamine reductase, imidazole acetate-4-monooxygenase, indole-2,3-dioxygenase, indole-3-acetaldehyde oxidase, isovaleryl-CoA dehydrogenase, kynurenine-3-monooxygenase, L-amino acid oxidase, L-aspartate oxidase, L-galactonolactone oxidase, L-glutamate oxidase, L-lactate dehydrogenase (cytochrome), lactate-2-monooxygenase, lathosterol oxidase, lathialciferin monooxygenase (demethylating), long-chain acyl-CoA dehydrogenase, lysine-2-monooxygenase, malate dehydrogenase (quinone), malate oxidase, mandelonitrile lyase, mellitate-3-monooxygenase, N-methyl-L-amino acid oxidase, NAD(P)+ Transhydrogenase (Si-specific), NAD(P)H dehydrogenase (quinone 1), NAD(P)H dehydrogenase (quinone), NADH peroxidase, NADPH dehydrogenase, NADPH dehydrogenase (quinone), NADPH-cytochrome-c2 reductase, NADPH-hemoprotein reductase, nicotinate dehydrogenase, nicotine dehydrogenase, nitrite reductase (NAD(P)H), nitrite reductase (NO formation), orcinol-2-monooxygenase, orotate reductase (NADH), orotate reductase (NADPH), oxalate oxidase, phenol-2-monooxygenase, phenylglyoxylate dehydrogenase (acylation), phthalate-4,5-dioxygenase, polyamine oxidase, proline dehydrogenase, putrescine oxidase, pyranose oxidase, pyridoxine-4-oxidase, pyridoxine-5-dehydrogenase, pyruvate dehydrogenase (cytochrome), pyruvate oxidase, pyruvate oxidase (CoA-acetylated), retinal dehydrogenase, rubredoxin-NAD(+) reductase, salicylate-1-monooxygenase, sarcosine dehydrogenase, short-chain acyl-CoA dehydrogenase, spermidine dehydrogenase, steroid-9α-monooxygenase, tartronate-semialdehyde synthase, taxifolin-8-monooxygenase, thiamine oxidase, trypanothione disulfide reductase, UDP-N-acetylmuramic acid dehydrogenase, or vanillyl alcohol oxidase.

[0132] Sensor membrane 13 and 14B-14D, in some embodiments of the disclosure, the sensor 500 or a portion of the sensor 500 includes an outer membrane 520 or 335. The outer membrane 520 or 335 overlies at least the working electrode 501 or 320 and the sensing element 322 or sensing region 509. Electrochemical sensors are often covered with an outer membrane 520 or 335 (e.g., a polymeric membrane) to provide stability to the sensing reagents (e.g., the analyte-specific enzyme 323 and the redox mediator 324), to limit mass transport, provide biocompatibility, and / or prevent fouling to the electrodes.

[0133] In some embodiments of the present disclosure, the membrane is made up of two components: a hydrophilic (water-loving) polymer and a crosslinker. The crosslinker bonds the polymer molecules together and to the sensing layer of the sensor. For analytes such as glucose, found in vivo at concentrations of about 5 mM, a flow-restricting membrane is necessary to prevent fouling of the electrodes. Flow-restricting sensor membranes are disclosed, for example, in U.S. Patent No. 6,932,894 to Mao et al., the entire contents of which are incorporated herein by reference.

[0134] For low concentrations of analytes, longer accumulation times (e.g., up to 30 minutes) and a flow-restricting membrane may be used. Alternatively, for low concentrations of analytes, a more permeable membrane may be used to maintain the natural flow of analytes to the sensing layer while having a membrane that enhances the biocompatibility of the sensor. For example, a hydrophilic membrane surface will not overwhelm the body's immune system, thereby reducing the risk of inflammation and other responses that may impair the performance of the sensor.

[0135] Sample Monitoring System Thus, these embodiments include analyte monitoring devices and systems that include an analyte sensor, at least a portion of which can be positioned beneath the surface of a user's skin, for in vivo detection of an analyte in a bodily fluid. Analyte monitoring systems are disclosed in Say et al. (U.S. Patent No. 6,134,461) and Hoss et al. (U.S. Patent Application Publication No. 2012 / 0150005), both of which are incorporated herein by reference in their entirety. Embodiments of the present disclosure include fully implantable analyte sensors, as well as analyte sensors in which only a portion of the sensor is positioned beneath the skin and another portion of the sensor is above the skin for communication with, for example, a sensor control device (which may include a transmitter), a receiver / display device, a transceiver device, a processor, etc. The sensor may be positioned, for example, beneath the skin of a user, for continuous or periodic monitoring of an analyte concentration in the interstitial fluid of the user. For purposes of this description, continuous monitoring and periodic monitoring are interchangeable unless otherwise noted. The sensor response may be correlated and / or converted to an analyte concentration in blood or other fluids. In certain embodiments, the analyte sensor may be placed in contact with the interstitial fluid to detect the analyte concentration, which may be used to estimate the analyte concentration in the user's bloodstream. The analyte sensor may be insertable into a vein, artery, or other part of the body that contains fluid. In some embodiments, the analyte sensor may be adapted to monitor the analyte concentration for a period of time, which may range from seconds, minutes, hours, days, weeks, months, or longer.

[0136] In some embodiments of the present disclosure, the analyte sensor may detect an analyte in vivo for an hour or more, e.g., several hours or more, days or more, three or more, five or more, seven or more days, weeks or more, or even a month or more. Future analyte concentrations may be predicted based on information obtained, such as the current analyte concentration at time t0, the rate of change of the analyte, etc. A predictive alarm may alert a user to a predicted analyte concentration of concern before the user's analyte concentration reaches the future predicted analyte concentration, thereby providing the user with an opportunity to take corrective action.

[0137] 15 illustrates a data monitoring and management system, such as an analyte monitoring system 400, according to certain embodiments of the present disclosure. Aspects of the embodiments of the present disclosure are further described primarily with respect to glucose monitoring devices and systems, and methods for detecting glucose. This description is for convenience only, and is in no way intended to limit the scope of the embodiments. It will be appreciated that the analyte monitoring system may be adapted to monitor various analytes disclosed herein simultaneously or at separate times.

[0138] Analytes that may be monitored include, but are not limited to, glucose, lactate, 3-hydroxybutyrate, cortisol, alcohol, pyruvate, glutamate, theophylline, acetylcholine, amylase, bilirubin, cholesterol, chorionic gonadotropin, glycosylated hemoglobin (HbA1c), creatine kinase (e.g., CK-MB), creatine, creatinine, DNA, fructosamine, glucose derivatives, glutamine, growth hormone, hormones, 3-hydroxybutyrate, ketones, ketone bodies, peroxide, prostate specific antigen, prothrombin, RNA, thyroid stimulating hormone, and troponin. Analytes may also include and be monitored, such as, for example, antibiotics (e.g., gentamicin, vancomycin, etc.), digitoxin, digoxin, drugs of abuse, theophylline, and warfarin. In some embodiments, more than one analyte may be monitored, and the analytes may be monitored simultaneously or at separate times.

[0139] The analyte monitoring system 400 includes an analyte sensor 401, a data processing device 402 connectable to the sensor 401, and a first receiving device 404. In some examples, the first receiving device 404 is adapted to communicate with the data processing device 402 via a communication link 403. In certain embodiments, the first receiving device 404 may further be adapted to transmit data to a data processing terminal 405 for evaluating, processing, or initializing the data received by the first receiving device 404. The data processing terminal 405 may be adapted to receive data directly from the data processing device 402 via a communication link 407, which may optionally be adapted for bidirectional communication. Additionally, the data processing device 402 may include a transmitting or transceiver device for transmitting and / or receiving data to the first receiving device 404 and / or the data processing terminal 405 and / or optionally the second receiving device 406.

[0140] 15, an optional second receiving device 406 may be operatively connected to the communication link 403 to receive data transmitted from the data processing device 402. The second receiving device 406 may be in communication with the first receiving device 404 and the data processing terminal 405. In some embodiments, the second receiving device 406 may be in two-way wireless communication with the first receiving device 404 and the data processing terminal 405. As described in more detail below, in some examples, the second receiving device 406 may be a receiving device with less functionality than the first receiving device 404, e.g., the second receiving device 406 may have a limited or minimal number of features compared to the first receiving device 404. Thus, the second receiving device 406 may include a smaller housing (including in one or more of all dimensions) or be embodied in a device including a watch, a wristband, a digital organizer, an mp3 player, a mobile phone, etc. Alternatively, the second receiving device 406 may have the same or substantially similar functionality as the first receiving device 404. The second receiving device 406 may include a connection portion adapted to mate with a bedside tabletop holder device for nighttime monitoring and / or a two-way communication device, for example. The tabletop holder may provide for recharging of the power source.

[0141] The embodiment of the analyte monitoring system 400 shown in FIG. 15 shows only one analyte sensor 401, data processing device 402, and data processing terminal 405. However, one of ordinary skill in the art will appreciate that the analyte monitoring system 400 may include more than one sensor 401 and / or more than one data processing device 402, and / or more than one data processing terminal 405. The multiple sensors may be placed within the user's body to monitor the analyte at the same time or at different times. In certain embodiments, analyte information obtained by a first sensor placed within the user's body may be used in comparison with analyte information obtained by a second sensor. This may be useful to confirm or validate analyte information obtained from one or both of the sensors. Such redundancy may be beneficial when the analyte information is to be considered in important treatment decisions. In certain embodiments, the first sensor may be used to calibrate the second sensor.

[0142] The analyte monitoring system 400 may be a continuous monitoring system, a semi-continuous monitoring system, or a discrete monitoring system. In a multiple component environment, each component may be uniquely identified by one or more of the other components in the system such that communication conflicts may be easily resolved between the various components in the analyte monitoring system 400. For example, a unique ID, a communication channel, etc. may be used.

[0143] In certain embodiments, the sensor 401 is physically located in or on the body of a user whose analyte concentration is being monitored. The sensor 401 may be adapted to at least periodically sample the user's analyte concentration and convert the sampled analyte concentration into a corresponding signal for transmission by the data processing device 402. The data processing device 402 may be coupled to the sensor 401, with either device being located in or on the user's body, with at least a portion of the analyte sensor 401 being located transcutaneously. The data processing device may include a fixation element (e.g., adhesive) for fixing the data processing device to the user's body. An attachment that is wearable on the user and engages with the data processing device 402 may be used. For example, the attachment may include an adhesive surface. The data processing device 402 performs data processing functions. Such functions may include, but are not limited to, filtering and encoding of data signals. Each of the data signals corresponds to a sampled analyte concentration of the user and is transmitted to a first receiving device 404 via a communication link 403. In some embodiments, the sensor 401 or data processing device 402, or a combination sensor / data processing device, may be entirely implantable beneath the surface of the user's skin.

[0144] In a particular embodiment, the first receiving device 404 may include an analog interface section including an RF receiver section and an antenna adapted to communicate with the data processing device 402 via the communication link 403, and a data processing section for processing data received from the data processing device 402. The processing may include data decoding, error detection and correction, generation of a data clock, data bit recovery, or any combination thereof.

[0145] In operation, the first receiving device 404 in certain embodiments is adapted to synchronize with the data processing device 402 to uniquely identify the data processing device 402, for example based on the identification information of the data processing device 402, and then periodically receive a signal transmitted from the data processing device 402 that is associated with the monitored analyte concentration detected by the sensor 401.

[0146] 15, the data processing terminal 405 may include a personal computer, a portable computer (laptop or handheld device (e.g., a personal digital assistant (PDA)), a telephone (including a mobile phone, e.g., a multimedia Internet enabled mobile phone, including an iPhone, Blackberry, or similar phone), an mp3 player (e.g., an iPOD, etc.), a pager, etc.), and / or a drug delivery device (e.g., an infusion device), each of which may be in data communication with a receiving device via a wired or wireless connection. The data processing terminal 405 may also be coupled to a data network (not shown) for storing, retrieving, updating, and / or analyzing data corresponding to the detected analyte concentration of the user.

[0147] The data processing terminal 405 may include a drug delivery device (e.g., an infusion device such as an insulin infusion pump) that may be adapted to administer a drug (e.g., insulin) to a user and may be in communication with the first receiving device 404 to receive, among other things, the measured analyte concentration. Alternatively, the first receiving device 404 may have an infusion device incorporated therein such that the first receiving device 404 administers the appropriate drug (e.g., insulin) to the user, for example to administer or modify a basal rate, among other things, to determine an appropriate bolus dose to administer based on the detected analyte concentration received from the data processing device 402. The infusion device may be an external device or an internal device, such as a device that may be entirely implanted in the user.

[0148] In certain embodiments, the data processing terminal 405 may include an infusion device, such as an insulin pump, and may be adapted to receive analyte signals from the data processing device 402, and may thus incorporate the functionality of the first receiving device 404, including data processing to manage the user's insulin therapy and analyte monitoring. In certain embodiments, the communication link 403 and one or more of the other communication interfaces shown in FIG. 15 may use one or more wireless communication protocols. Such protocols include, but are not limited to, RF communication protocols, infrared communication protocols, Bluetooth-enabled communication protocols, 802.11x wireless communication protocols, or equivalent wireless communication protocols that allow for secure wireless communication of several units (e.g., per Health Insurance Portability and Accountability Act (HIPPA) requirements) while avoiding potential data collisions and interference.

[0149] In further embodiments, the data processing device 402 and / or the first receiving device 404 and / or the second receiving device 406 and / or the data processing terminal (infusion device) 405 may be adapted to receive the analyte values ​​wirelessly, for example from a blood analyte meter, via a communications link. In further embodiments, a user operating or using the analyte monitoring system 400 (FIG. 15) may manually input the analyte values ​​using, for example, a user interface (e.g., a keyboard, keypad, voice commands, etc.) associated with one or more of the data processing device 402, the first receiving device 404, the second receiving device 406, or the data processing terminal (infusion device) 405.

[0150] The sensors (e.g., enzymatic biosensors) disclosed herein that measure low nanomolar concentrations of analytes may be used in an in-vivo monitoring system that is placed in the body of a user (e.g., a human subject) to contact the user's bodily fluids and sense the concentration of one or more analytes therein. The in-vivo monitoring system may include one or more readers that receive sensed analyte data from a sensor control device. These readers may process and / or display the sensed analyte data or sensor data to a user in any number of forms.

[0151] 16, in some embodiments, the reader 120 may be a mobile communication device such as a dedicated reader (adapted to communicate with the sensor control device 102 (FIG. 17) and optionally a computer system, but not mobile phone communication capabilities) or may be a mobile phone (including, but not limited to, a Wi-Fi or Internet enabled smartphone, tablet, or personal digital assistant (PDA)). Examples of smartphones include mobile phones based on the Windows® operating system, Android® operating system, iPhone® operating system, Palm®, WebOS®, Blackberry® operating system, or Symbian® operating system with data network connectivity for Internet connectivity and / or data communication over a local area network (LAN).

[0152] The reader 120 may also be configured as a mobile smart wearable electronic assembly, such as an optical assembly worn on or adjacent to a user's eye (e.g., a mobile communication device, such as Google Glasses, or other smart glasses). The optical assembly may have a transparent display that displays information to the user regarding the user's analyte concentration, while allowing the user to see through the display while simultaneously obscuring the user's overall field of vision as little as possible. The optical assembly may also be capable of wireless communication similar to a smart phone. Other examples of wearable electronic devices include devices worn around or near a user's wrist (e.g., a watch, etc.), neck (e.g., a necklace), head (e.g., a headband, hat, etc.), chest, etc.

[0153] FIG. 16 is a block diagram of an exemplary embodiment of a reader 120 configured as a smartphone. Here, the reader 120 includes an input 121, a display 122, and a processing circuit 206. The processing circuit 206 may include one or more processors, microprocessors, controllers, and / or microcontrollers, each of which may be individual chips or may be distributed among several different chips (and portions thereof). Here, the processing circuit 206 includes a communication processor 202 having on-board memory 203 and an application processor 204 having on-board memory 205. The reader 120 further includes an RF communication circuit 208 connected to an RF antenna 209, a memory 210, a multifunction circuit 212 having one or more associated antennas 214, a power source 216, a power management circuit 218, and a clock 219. FIG. 16 omits the representation of typical hardware and functions present in a smartphone. Other hardware and functions (e.g., codecs, drivers, glue logic) may also be included, as would be apparent to one skilled in the art.

[0154] 16, the communications processor 202 may interface with RF communications circuitry 208 to perform analog-to-digital conversion, encoding and decoding, digital signal processing, and other functions to facilitate conversion of voice, video, and data signals into a suitable format (e.g., in-phase and quadrature) for presentation to the RF communications circuitry 208, which may then transmit the signals wirelessly. The communications processor 202 may also interface with the RF communications circuitry 208 to perform the inverse functions necessary to receive wirelessly transmitted signals and convert them into digital data, voice, and video. The RF communications circuitry 208 may include transmit and receive sections (e.g., integrated as a transceiver section) and associated encoder logic.

[0155] 16, the application processor 204 may be adapted to execute the operating system and any software applications present on the reader 120, process video and images, and perform other functions not related to processing communications transmitted and received via the RF antenna 209. The smartphone operating system operates in conjunction with the applications present on the reader 120. Any number of applications (also known as "user interface applications") may be running on the reader 120 at any one time, including one or more applications related to a diabetes monitoring regimen, as well as other commonly used applications unrelated to such regimen (e.g., email, calendar, weather, support, games, etc.). For example, data received by the reader indicative of sensed analyte concentrations and in vitro blood analyte measurements may be securely communicated to the user interface applications present in the memory 210 of the reader 120. Such communication may be performed securely, for example, using mobile application containerization or wrapping techniques.

[0156] The memory 210 may be shared by one or more of the various functional units present in the reader 120 or may be distributed among two or more of them (e.g., as separate memories present in different chips). The memory 210 may also be a separate chip itself. The memories 203, 205, and 210 may be non-transient and may be volatile (e.g., RAM, etc.) and / or non-volatile (e.g., ROM, Flash memory, F-RAM, etc.) memories. The multi-function circuit 212 may be implemented as one or more chips and / or components (e.g., transmitters, receivers, transceivers, and / or other communication circuitry) that may perform other functions, such as, for example, local wireless communication with the sensor controller 102 and determining the geographic location of the reader 120 (e.g., Global Positioning System (GPS) hardware) in an appropriate protocol (e.g., Wi-Fi, Bluetooth, Bluetooth Low Energy, Near Field Communication (NFC), Radio Frequency Identification (RFID), proprietary protocols, and other protocols). One or more other antennas 214 may be associated with the functional circuitry 212 as needed to operate with various protocols and circuits.

[0157] The power source 216 may include one or more batteries. These batteries may be rechargeable or disposable. The power management circuitry 218 may coordinate battery charging, power supply monitoring, and may provide power up and DC conversion.

[0158] The reader 120 may also include or be incorporated into a drug (e.g., insulin, etc.) delivery device, for example, sharing a common housing. Examples of such drug delivery devices include drug pumps (e.g., wearable pumps that deliver basal and bolus doses of insulin) with a cannula that can be placed in the body to infuse for hours or days. When combined with a drug pump, the reader 120 may include a container for storing the drug, a pump connectable to a delivery tube, and an infusion cannula. The pump may move the drug from the container through a tube and into the diabetic patient's body through a cannula inserted in the body. Other examples of drug delivery devices that may be included (or incorporated) in the reader 120 include portable injection devices (e.g., insulin pens) that are inserted into the skin only for each delivery and then removed. When combined with a portable injection device, the reader 120 may include a needle, a cartridge that carries the drug, an interface to adjust the amount of drug delivered, and an actuator to trigger the injection. The device may be used repeatedly until the drug is exhausted. Once the drug is depleted, the combined device may be discarded or the cartridge may be replaced with a new one. At this point, the combined device may be reused repeatedly. The needle may be replaced after each injection.

[0159] The combined device may function as part of a closed loop system (e.g., an artificial pancreas system that requires no user intervention to operate) or a semi-closed loop system (e.g., an insulin loop system that requires minimal user intervention to operate, such as confirming dosage changes). For example, the diabetic patient's analyte concentration may be automatically and repeatedly monitored by the sensor controller 102. The sensor controller 102 may then communicate the monitored analyte concentration to the reader 120 to automatically determine and deliver to the diabetic patient an appropriate drug dosage to adjust the diabetic patient's analyte concentration. Software instructions for controlling the pump and the amount of insulin delivered may be stored in memory of the reader 120 and executed by the reader's processing circuitry. These instructions may also cause calculation of drug delivery amount and duration (e.g., bolus and / or basal doses) based on analyte concentration measurements obtained directly or indirectly from the sensor controller 102. In some embodiments, the sensor controller 102 may determine and communicate the drug dosage to the reader 120.

[0160] FIG. 17 is a block diagram illustrating an exemplary embodiment of a sensor controller 102. The sensor controller 102 includes an analyte sensor 104 and sensor electronics 250 (including analyte monitoring circuitry) that may include a significant portion of processing capability to render final result data in a form suitable for display to a user. In FIG. 17, a single semiconductor chip 251 is shown, but this may be a custom application specific integrated circuit (ASIC). Within the ASIC 251, certain higher level functional units are shown, including an analog front end (AFE) 252, power management (i.e., control) circuitry 254, a processor 256, and communication circuitry 258 (which may be implemented as a transmitter, receiver, transceiver, passive circuitry, or according to a communication protocol). In this embodiment, both the AFE 252 and the processor 256 are used as analyte monitoring circuitry, but in other embodiments, either circuitry may perform the analyte monitoring function. The processor 256 may include one or more processors, microprocessors, controllers, and / or microcontrollers, each of which may be separate chips and / or distributed among several different chips (and portions thereof).

[0161] The memory 253 may also be included in the ASIC 251, shared by the various functional units present in the ASIC 251, or distributed between two or more of them. The memory 253 may also be a separate chip. The memory 253 is non-transient and may be volatile and / or non-volatile memory. In this embodiment, the ASIC 251 is connected to a power source 260. The power source 260 may be a coin cell battery or the like. The AFE 252 is connected to the in-vivo analyte sensor 104 to receive measurement data therefrom and outputs the data in digital form to the processor 256. The processor 256 then processes the data in any suitable manner, depending on the embodiment. The data may be provided to the communication circuitry 258 and transmitted, for example, by the antenna 261 to the reader 120. At the reader 120, minimal further processing is required by the resident software application to display the data. The antenna 261 may be configured as required by the application and communication protocol. The antenna 261 may be, for example, a printed circuit board (PCB) wired antenna, a ceramic antenna, or a separate metal antenna. The antenna 261 may be configured as a monopole antenna, a dipole antenna, an F-antenna, a loop antenna, and other antennas.

[0162] Information may be communicated from the sensor controller 102 to a second device (e.g., the reader 120) at the initiative of the sensor controller 102 or the reader 120. For example, information may be communicated by the sensor controller 102 automatically and / or repeatedly (e.g., continuously) when analyte information is available, or on a scheduled basis (e.g., about every minute, about every five minutes, about every ten minutes, etc.). In the latter case, the information may be stored or recorded in the memory of the sensor controller 102 for later communication. Information may be transmitted from the sensor controller 102 in response to receiving a request by the second device. The request may be an automated request, e.g., a request transmitted by the second device on a scheduled basis, or may be a request generated at the initiative of a user (e.g., an ad-hoc request or a manual request). In some embodiments, a manual request for data is referred to as a "scan" of the sensor controller 102 or an "on-demand" transmission of data from the device 102. In some embodiments, the second device may send a polling signal or data packet to the sensor controller 102, which may treat each poll (or multiple polls occurring at a particular time interval) as a request for data and, if data is available, transmit the data to the second device. In many embodiments, communications between the sensor controller 102 and the second device are secure (e.g., encrypted communications and / or communications between authenticated devices), but in some embodiments, data may be sent from the sensor controller 102 in an insecure manner, for example, as a broadcast to all viewing devices within range.

[0163] Different types and / or forms and / or amounts of information may be transmitted as part of each communication, including, but not limited to, one or more current sensor measurements (e.g., most recently obtained analyte concentration information corresponding in time to the time a reading was initiated), the rate of change of measurements taken over a given period of time, the velocity of the rate of change of measurements (rate of rise of the rate of change), or measurement history information corresponding to measurement information taken prior to a given reading and stored in the memory of the sensor controller 102.

[0164] Any or all of the real-time information, historical information, rate of change information, and rate of change (e.g., rate of increase or decrease) information may be transmitted to the reader 120 in a given communication or transmission. In certain embodiments, the type and / or format and / or amount of information transmitted to the reader 120 may be preprogrammed and / or non-alterable (e.g., preset at time of manufacture), or may not be preprogrammed and / or non-alterable such that it is selectable (e.g., by turning on the system) and / or may be altered one or more times in the field. Thus, in certain embodiments, the reader 120 may output the current (real-time) sensor-derived analyte value (e.g., in the form of a numerical value), the current rate of analyte change (e.g., in the form of an indication of the rate of analyte change, such as an arrow pointing in the direction of the current rate), and historical analyte trend data (e.g., in the form of a graph) based on sensor readings acquired and stored by the memory of the sensor control device 102. Skin surface or sensor temperature readings or measurements may also be collected by the optional temperature sensor 257. These readings or measurements may be communicated (individually or as aggregate measurements over time) from the sensor controller 102 to other devices (e.g., a reader or reader 120). However, instead of or in addition to actually displaying a temperature measurement to a user, the temperature readings or measurements may be used in conjunction with software routines executed by the reader 120 to correct or compensate the analyte measurement output to the user.

[0165] The following examples are provided for illustrative purposes only and are not intended to limit the scope or content of the present application. EXAMPLES

[0166] Example 1: Sensitivity calculations for accumulation mode detection using polymer-coated sensors and long accumulation times Figure 5 shows the calibration curves obtained by amperometric and accumulation mode sensing using polymer-coated glucose sensors at glucose concentrations from 0 to 500 μM. Each calibration curve is the average response of four sensors. However, unlike amperometric measurements, the sensitivity of the sensor by accumulation mode sensing can be easily adjusted by varying the accumulation time. Increasing the accumulation time from 1 to 10 min increases the sensitivity of the sensor by approximately 10 times for both peak height and peak area measurements. The sensitivity of each calibration curve shown in Figure 5 was calculated from the slope of the linear regression with the tabulated data shown in Table 1.

[0167] [Table 1]

[0168] Since the peak height and amperometric measurements are made in the same units, their sensitivities may be compared directly. Using data from a flow-limiting membrane sensor as shown in Figure 5, the ratio of sensitivity in accumulation mode to amperometric sensitivity for equivalent sensor conditions (i.e., the fold increase) was calculated using the data tabulation shown in Table 2. As shown, for an accumulation time of 1 minute, the sensitivity of the sensor is 2 times higher using accumulation mode sensing compared to amperometric. Thus, by increasing the accumulation time to 10 minutes, the difference in sensitivity increases by 15 times.

[0169] [Table 2]

[0170] Example 2: Increasing the frequency and adding carbon nanotubes to optimize accumulation mode signals for sensitive detection Figure 7 shows the accumulation mode detection of 200 nM glucose at two different signal filtering frequencies, 0.032 Hz and 3.2 Hz. As shown, the detection peak is much steeper and the peak height is larger when using the high frequency filter. However, the area under the two curves remains unchanged. This indicates that a high frequency filter is ideal to optimize the signal magnitude when using peak height measurements. Notably, it was found that changing the filtering frequency from 0.032 Hz to 3.2 Hz increased the peak height signal by 2-3 times. Furthermore, when the filtering frequency was higher than 3.2 Hz, the signal noise became too large to allow accurate measurement of either the amperometric current or the accumulation peak characteristics (peak height and peak area).

[0171] As a mechanistic means of enhancing the accumulation mode signal, carbon nanotubes (CNTs) were added to make the more uniformly deposited sensing reagent more electrically conductive, thereby enhancing the kinetics of the redox mediator-mediated oxidation process. Due to the increased kinetics, the accumulation mode current spikes had a larger peak height. Figure 8A shows micrographs of the deposited and cured glucose sensing reagent with and without CNTs. As shown, the sensing reagent with CNTs was more uniformly deposited, while the sensing reagent without CNTs exhibited a pronounced "coffee ring effect." The addition of CNTs to the sensing reagent was found to enhance the peak height signal by 5-6 times.

[0172] FIG. 8B also shows experimental results demonstrating the effect on sensor sensitivity of both signal filtering frequency and the addition of CNTs to the sensing reagent. The experiment used amperometric and accumulation mode sensing, with measurements by peak height and peak area using an exemplary glucose sensor as shown, with glucose concentrations from 0 to 200 nM. Four sensors of both types (with and without CNTs in the sensing reagent) were tested. Each calibration curve is the average response of the four sensors shown. The accumulation time used for each accumulation mode detection was 10 minutes. Two consecutive measurements were made at each glucose concentration, one using a filtering frequency of 0.032 Hz and the other using a filtering frequency of 3.2 Hz.

[0173] The sensitivity of each calibration curve shown in Figure 8B was calculated as the slope of the linear regression. The tabulated data are shown in Table 3. As is self-evident, the amperometric sensor sensitivity changes little with filtering frequency and the presence or absence of CNTs, being less than 0.0003 nA / nM for all conditions. For accumulation mode measurements using peak area, the sensor sensitivity does not change with filtering frequency, but increases slightly when CNTs are added to the sensing reagent. The largest change in sensor sensitivity was observed for accumulation mode measurements using peak height. Both filtering frequency and the addition of CNTs to the sensing reagent increase the sensor sensitivity. Increasing the filtering frequency from 0.032 Hz to 3.2 Hz increases the sensitivity by about 2.5 times, while adding CNTs to the sensing reagent increases the sensitivity by about 5.5 times. Furthermore, when the filtering frequency is combined with the addition of CNTs, the sensitivity of the accumulation mode measurement increases by about 14 times.

[0174] [Table 3]

[0175] The peak height and amperometric measurements are measured in the same units, so their sensitivities may be compared directly. Table 4 shows the ratio of the accumulation mode sensitivity to the amperometric sensitivity under equivalent sensor conditions. As shown, even with a filtering frequency of 0.032 Hz and no CNTs in the sensing reagent, the sensor sensitivity is 30 times higher using accumulation mode sensing compared to amperometric measurements. Therefore, by increasing the filtering frequency and adding CNTs to the sensing reagent, the accumulation mode peak height can be optimized, increasing the sensitivity difference to nearly 400 times.

[0176] [Table 4]

[0177] Example 3: Comparison of sensitivity, detection limit, and linear range in amperometric and accumulation mode sensing with carbon nanotube loading at an accumulation time of 30 min and a signal frequency of 3.2 Hz. As shown in FIG. 9B, the current associated with the amperometric measurements is very low (<50 pA) and loses linearity below 100 nM, while the signal of the sensing in accumulation mode is very high and remains well linear below 100 nM. Table 5 below shows the sensitivity, lower limit of detection (LOD) (calculated as 3σ / slope using standard method 1), and linear detection range associated with these measurements as disclosed in Example 5. Standard method 1 is disclosed in Mocak et al., Pure Appl. Chem. 1997, 69:297-328, the entire contents of which are incorporated herein by reference. In particular, standard method 1 is a method for calculating the LOD as "3σ / slope", where "σ" is the standard deviation of the blank test, and "slope" is the slope of the calibration curve.

[0178] [Table 5]

[0179] Example 4: Analysis of background signals With reference to Figures 9A and 9B, a negative (cathodic) background signal is observed when sensing in buffer solution exposed to air. Without being limited by any theory, it is believed that an oxygen reduction reaction is responsible for this negative background. Specifically, the osmium redox mediator and CNTs can catalyze the oxygen reduction reaction, which results in the oxidation of the osmium mediator and the breakdown of the circuit during the accumulation period, resulting in the formation of Os3 + When the circuit is reconnected, Os3 + This accumulation of is reduced, resulting in a cathodic peak. To test this hypothesis, an exemplary glucose sensor was tested in 100 mM phosphate buffer without glucose at atmospheric conditions and at conditions purged with oxygen (e.g., by air bubbling). FIG. 10A shows the accumulation mode signal obtained by a representative sensor at 2, 5, and 10 minutes of accumulation time at atmospheric and oxygen purged conditions as shown. As can be observed, the signal is cathodic peak at atmospheric conditions and a smaller anodic peak at oxygen purged conditions. The median (average) signal of the four sensors is depicted in FIG. 10B. As shown, the amperometric signal is observed as slightly negative at atmospheric conditions and slightly positive at oxygen purged conditions. The results of this experiment revealed that the negative background is due to Os-catalyzed oxygen reduction.

[0180] Example 5: Linear detection range To determine the linear detection range of sensing in accumulation mode, the calibration experiments shown in Figures 9A and 9B were performed up to a glucose concentration of 200 μM. The resulting amperometric and accumulation mode calibration curves are shown in Figure 11. The first-order best-fit lines determined for concentrations from 0 to 200 nM were predicted for higher concentrations. As can be seen, the amperometric signal was linear at least up to 100 μM. On the other hand, the accumulation mode signal was linear up to 2 to 5 μM and plateaued at higher concentrations. This was expected since the Os redox mediator has a finite charge storage capacity. For the sensor used in this experiment, this capacity appears to be up to about 5000 nC. It should be noted that the linear range of sensing in accumulation mode may shift to higher concentrations if shorter accumulation times are used. For the data presented here, high sensitivity was obtained using relatively long (e.g., 30 min) accumulation times.

[0181] Example 6: Materials Carbon sensors screen-printed on PET substrates were obtained from Steven Label, Inc. (Santa Fe Springs, CA). The active area of ​​the working electrode was approximately 0.1 mm2 in deposition area. 2 The glucose oxidation catalyst, glucose oxidase (GO), was identified. x The proprietary redox polymers used for the interconnects and the proprietary flow-limiting membrane polymers were synthesized according to published procedures and obtained from Nanosyn, Inc. (Santa Rosa, CA) and Regis Technologies, Inc. (Morton Grove, IL), respectively. Glucose oxidase (GO) from Aspergillus sp. II was used. x, EC1.1.3.4, activity: 130 U / mg) were obtained from Toyobo Co., Ltd. (Osaka, Japan). Polyethylene glycol (400) diglycidyl ether (PEGDGE400) and glyceryl triglycidyl ether were obtained from Polysciences, Inc. (Warrington, PA). Multi-walled carbon nanotubes (CNTs, outer diameter: 20–40 nm, length: 10–20 μm) were obtained from MK Nano (Mississauga, Ontario, Canada). Common chemicals used for glucose and buffer were obtained from Sigma-Aldrich (St. Louis, MO). Water (18.0 MΩ·cm) was obtained from a Thermo Scientific Barnstead E-Pure ultrapure water purification system. -1 All aqueous solutions were prepared using >100% deionized water.

[0182] Example 7: Preparation of a sensor Two different glucose sensing reagents were used, one with and one without CNTs. The non-CNT reagent was prepared as follows: 4% (w / v) redox polymer, 8.08% (w / v) GO. x Three solutions of PEGDGE400 at 8.08% (w / v) and 4% redox polymer were prepared in 10 mM 4-(2-hydroxyethyl)piperazine-1-ethanesulfonic acid (HEPES) buffer (pH: 8). These three solutions were mixed in a ratio of 3.04:5.1:1.86 to obtain the glucose sensing reagent. To prepare the glucose sensing reagent containing CNTs, the above procedure was followed, except that the 4% redox polymer solution and the 8.08% PEGDGE400 solution were prepared in 5% (w / v) aqueous CNT solution instead of 10 mM HEPES solution. After preparation, 15 nL of the glucose sensing reagent was delivered to the carbon working electrode of the sensor by a microsyringe (Hamilton Co.). The active area of ​​each working electrode was defined by the area of ​​the delivered drop of sensing reagent. This area was typically 0.1 mm 2After application of the sensing reagent, the sensor was cured at 25° C. and 60% relative humidity for at least 12 hours. For the sensor used in the experiment shown in FIG. 5, the sensor was coated with a flow-restricting polymer outer membrane. This membrane consisted of a 4:1 volumetric mixture of 14% (w / v) membrane polymer in 80 / 20 ethanol / water and 3.5% (w / v) glyceryl triglycidyl ether, applied by dip coating as described above in Liu et al., Anal. Chem. 2012, 84:3403-3409, the entire contents of which are incorporated herein by reference.

[0183] Example 8: Electrochemical measurements Unless otherwise noted, all electrochemical measurements were performed using a suitable three electrode cell with the glucose sensor as the working electrode, an Ag / AgCl reference electrode (3 M in KCl solution; Bioanalytical Systems, Inc.), and a screen-printed carbon counter electrode. The sensor current versus time was measured through accumulation mode experiments using a potentiostat. For accumulation mode measurements, the working electrode was electrically disconnected from the potentiostat for a set amount of time (accumulation time). The working electrode was then reconnected to the circuit. Figure 2 shows a schematic of the electrodes. When the sensor working electrode was electrically connected, it was +40 mV. For the experiments shown in Figures 3A-3D, 4A-4B, 5, and 6A-6H, a BASi Petit Ampere potentiostat (model: LC-3D; Bioanalytical Systems, Inc., West Lafayette, IL) was used to measure the current. A 0.03 Hz filter was used with a 0.5 s sampling interval. Current signals were recorded in-house using LabView (National Instruments) software. For all other experiments, increased time resolution was desired. Therefore, a potentiostat with high time resolution was used (Model: 1030C; CH Instruments, Inc., Austin, TX). The potentiostat used a 3.2 Hz filter with a 0.1 s sampling interval, except for those shown in Figures 7 and 8B. As indicated, for these experiments, the potentiostat used a 0.1 s sampling interval and either a 3.2 Hz filter or a 0.032 Hz filter. The signals were recorded using software provided by the manufacturer. Measurements of current peak area, peak height, and amperometric current versus time were performed using Graphpad Prism 6 software. All experiments were performed in 100 mM PBS buffer (pH = 7.4, 100 mM NaCl) at 33 °C.

[0184] As disclosed and illustrated throughout this specification, accumulation mode sensing according to embodiments of the present disclosure may be used to provide superior detection than amperometric measurements at low analyte concentrations.

[0185] While the present disclosure has been illustrated and described with reference to specific exemplary embodiments, it would be apparent to those skilled in the art that various modifications and changes can be made to the described embodiments without departing from the spirit and scope of the present disclosure, as defined in the appended claims.

Claims

1. A continuous analyte monitoring system including an in-vivo analyte sensor and sensor electronics, the in vivo analyte sensor includes a distal portion configured to be transcutaneously placed beneath a skin surface of a subject, the distal portion including a working electrode, a counter electrode, and a reference electrode, and a proximal portion configured to be electrically connected to the sensor electronics on the skin surface of the subject; the sensor electronics including one or more processors coupled to a communications circuit; and a non-transitory memory; the non-transitory memory storing instructions that, when executed by the one or more processing devices, cause the one or more processing devices to generate data indicative of an analyte concentration in the subject based on one or more electrical signals received from the in vivo analyte sensor and wirelessly transmit the data indicative of the analyte concentration in the subject via a communication circuit to a display device in accordance with a Bluetooth communication protocol; the working electrode comprises a sensing element comprising an analyte-specific enzyme and a redox mediator, the sensing element being configured to accumulate charge derived from an analyte reacting with the analyte-specific enzyme over a period of time and to connect to the working electrode after the period of time and measure a signal from the accumulated charge; The system wherein the redox mediator comprises a redox species selected from osmium, ruthenium, iron, cobalt, and compounds or complexes thereof, bound to a polymer selected from poly(vinylpyridine), poly(thiophene), poly(aniline), poly(pyrrole), and poly(acetylene).

2. The system of claim 1, wherein the materials of the working electrode, counter electrode and reference electrode are selected from aluminum, carbon, cobalt, copper, gallium, gold, indium, iridium, iron, lead, magnesium, mercury (as mercury alloys), nickel, niobium, osmium, palladium, platinum, rhenium, rhodium, selenium, silicon, silver, tantalum, tin, titanium, tungsten, uranium, vanadium, zinc, zirconium, mixtures thereof, alloys thereof, oxides thereof and metal compounds thereof.

3. The system described in claim 1, wherein measuring the signal from the accumulated charge includes filtering the signal at a frequency of 0.032 to 3.2 hertz (Hz).

4. The working electrode includes a first conductive layer and a sensing area on the first conductive layer; the reference electrode includes a second conductive layer and a second conductive material on the second conductive layer; and the counter electrode includes a third conductive layer; The system of claim 1 , wherein the sensing area includes the sensing element.

5. The in vivo analyte sensor further comprises: a first insulating layer over at least a portion of the first conductive layer; and a second insulating layer over at least a portion of the second conductive layer; the second conductive layer overlies at least a portion of the first insulating layer; The system of claim 4 , wherein the third conductive layer overlies at least a portion of the second insulating layer.

6. The system described in claim 5, wherein the first insulating layer and the second insulating layer are each independently a dielectric layer.

7. The system of claim 1, wherein an outer membrane covers at least the sensing element.

8. The system of claim 1, wherein the reference electrode and the counter electrode are a single counter / reference electrode.

9. The system further includes a display device, the display device comprising: wireless communication circuitry of a display device, the wireless communication circuitry configured to receive data indicative of an analyte concentration transmitted from the sensor electronics; and and one or more processing units of the display device coupled to a non-transitory memory of the display device, wherein the non-transitory memory of the display device is configured to store an analyte monitoring software application that, when executed by the one or more processing units of the display device, causes the one or more processing units of the display device to: displaying an indicative display of a trend based on the received data indicative of the analyte concentration; displaying an analyte concentration based on the received data indicative of the analyte concentration; generating an alarm based on the received data indicative of the analyte concentration; and / or Controlling a drug delivery device; The system of claim 1 .

10. The system of claim 1, wherein the sensing element further comprises a carbon nanotube.

11. The system of claim 1, wherein the analyte is selected from the group consisting of cortisol, glucose, lactate, 3-hydroxybutyrate, alcohol, pyruvate, glutamate, theophylline, and creatinine.

12. The system of claim 1, wherein the analyte-specific enzyme is selected from the group consisting of nicotinamide adenine dinucleotide (NAD)-dependent dehydrogenase, flavin adenine dinucleotide (FAD)-dependent oxidase, and flavin mononucleotide (FMN)-dependent oxidase.

13. The system of claim 1, wherein the sample-specific enzyme is selected from the group consisting of 11β-hydroxysteroid dehydrogenase type 2 (11β-HSD-2), glucose oxidase, NAD-glucose dehydrogenase, FAD-glucose dehydrogenase, lactate oxidase, NAD-lactate dehydrogenase, NAD-alcohol dehydrogenase, pyruvate oxidase, NAD-glutamate dehydrogenase, and xanthine oxidase.

14. The system of claim 1, wherein the redox mediator is a redox polymer immobilized on the working electrode.

15. The system of claim 1, wherein the sample has a concentration of 4.7 nanomolar or greater.

16. A method of sensing an analyte using the continuous analyte monitoring system of any one of claims 1 to 15, comprising: providing said working electrode having a sensing element comprising an analyte-specific enzyme and a redox mediator; providing the working electrode to the analyte; storing a charge derived from the analyte which reacts with the analyte-specific enzyme and the redox mediator; connecting the working electrode to a circuit after a predetermined period of time; and measuring a signal from the accumulated charge.

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