Methods for detection of an analyte using an electrochemical biosensor and system for detection of an analyte

Abstract

The present invention relates to a method for detecting an analyte using a sensor that has an operating electrode, wherein the method includes providing the operating electrode with an analyte-specific enzyme and a redox mediator, providing the operating electrode to the analyte, accumulating the charge derived from the analyte reacting with the analyte-specific enzyme and the redox mediator for a specified period of time, connecting the operating electrode to the circuit after the specified period of time, and measuring the signal of the accumulated charge.

Description

1 / 76 “METHODS FOR THE DETECTION OF AN ANALYTE USING AN ELECTROCHEMICAL BIOSENSOR AND SYSTEM FOR THE DETECTION OF A "ANALITO" REFERENCE TO RELATED REQUEST(S)

[0001] This patent application claims the benefits of U.S. Patent Application Serial No. 16 / 024,353, filed June 29, 2018; U.S. Provisional Patent Application Serial No. 62 / 527,981, filed June 30, 2017; U.S. Provisional Patent Application Serial No. 62 / 544,692, filed August 11, 2017; and U.S. Provisional Patent Application Serial No. 62 / 545,252, filed August 14, 2017, the full contents of which are expressly incorporated herein by reference. DECLARATION REGARDING RESEARCH OR DEVELOPMENT SPONSORED BY THE FEDERAL GOVERNMENT

[0002] The present invention was developed with government support under contract no. HDTRA-1-16-C-0048 granted by the Defense Threat Reduction Agency. The government has certain rights in the invention. FIELD

[0003] The embodiments of the present invention relate to analyte detection by using electrochemical enzymatic biosensors. For example, the embodiments of the present invention relate to a method and an enzymatic biosensor that allows the detection of low analyte concentrations, allowing an accumulation of the analyte in the biosensor. BACKGROUND

[0004] Enzymatic biosensors that utilize enzymes associated with a transducer as a biorecognition element for a target analyte have been developed and are used. Although many different signal transduction methods are used, the most frequently used has been the electrochemical method. The Petition 870250049834, dated 06 / 13 / 2025, p. 10 / 97 2 / 76 Electrochemical biosensors allow the biological event (e.g., analyte detection) to be directly converted into an electrical signal, eliminating the need for complex instrumentation and thus giving electrochemical biosensors desirable characteristics in terms of size, cost, and portability. Among the electrochemical techniques used for signal transduction, amperometry is frequently used. In an amperometric measurement, the operating electrode of the sensor is maintained at a constant potential (voltage) while the current flowing through the sensor is measured. The sensor is designed in such a way that the current is dependent on the analyte concentration.

[0005] An example of an enzymatic biosensor that uses amperometry is the continuous glucose sensor, which is a transportable in vivo device designed to provide frequent measurements of blood glucose concentration to the user. These devices utilize a glucose oxidoreductase enzyme, such as glucose oxidase (GOx), immobilized on an operating electrode as the glucose sensing element. Electrons are first passed from glucose to the enzyme via enzymatic oxidation, and then to the operating electrode via a redox mediator, such as an oxygen (O2) or osmium (Os)-containing redox polymer. While amperometry has proven viable for measuring analytes such as glucose, which is present at relatively high physiological concentrations (at or above 5 millimolar (mM)), it may not be appropriate for measuring analytes present at lower concentrations. SUMMARY

[0006] Aspects of the embodiments of the present invention relate to the detection of low concentrations (e.g., 5 mM or less, 1 nanomolar (nM) at 5 mM, or 4.7 nM at 5 mM) of the analyte by allowing an accumulation of analyte in an enzymatic biosensor. Petition 870250049834, dated 06 / 13 / 2025, p. 11 / 97 3 / 76

[0007] In some embodiments of the present invention, it is a method for detecting an analyte that makes use of a sensor having an operating electrode, wherein the method includes providing the operating electrode with an analyte-specific enzyme and a redox mediator, providing the operating electrode to the analyte, accumulating the charge derived from the analyte reacting with the analyte-specific enzyme and the redox mediator for a stipulated period of time, connecting the operating electrode to a circuit after the stipulated period of time, and measuring a signal of the accumulated charge.

[0008] In some embodiments of the present invention, prior to providing the working electrode to an analyte, the method includes connecting the working electrode to the circuit, and prior to providing the working electrode to the analyte, the method includes disconnecting the working electrode from the circuit.

[0009] In some embodiments of the present invention, the working electrode is connected to the circuit before the working electrode is provided to the analyte, and the method includes disconnecting the working electrode from the circuit before the working electrode is provided to the analyte.

[0010] In some embodiments of the present invention, the sensor is an enzymatic electrochemical biosensor.

[0011] In some embodiments of the present invention, the redox mediator is an immobilized redox polymer.

[0012] In some embodiments of the present invention, the immobilized redox polymer includes a redox species and a polymer, wherein the redox species is selected from a polymer containing osmium (Os), ruthenium (Ru), iron (Fe) or cobalt (Co), and the polymer is selected from poly(vinylpyridine), poly(thiophene), poly(aniline), poly(pyrrole), or poly(acetylene).

[0013] In some embodiments of the present invention, the immobilized redox polymer is a poly(vinylpyridine) containing Os. Petition 870250049834, dated 06 / 13 / 2025, p. 12 / 97 4 / 76

[0014] In some embodiments of the present invention, the analyte is selected from cortisol, glucose, lactate, 3-hydroxybutyrate, alcohol, pyruvate, glutamate, theophylline or creatinine.

[0015] In some embodiments of the present invention, 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] In some embodiments of the present invention, the analyte-specific enzyme is selected from 11β-hydroxysteroid dehydrogenase type 2 (11β-HSA-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] In some embodiments of the present invention, charge accumulation includes the accumulation of electrons.

[0018] In some embodiments of the present invention, the sensor is placed subcutaneously in an individual.

[0019] In some embodiments of the present invention, the analyte is at a concentration as low as 4.7 nanomolars (nM).

[0020] In some embodiments of the present invention, the stipulated time period varies from 60 seconds to 30 minutes. In some embodiments, the stipulated time period varies from 120 seconds to 30 minutes. In some embodiments, the stipulated time period varies from 120 seconds to 10 minutes.

[0021] In some embodiments of the present invention, the sensor includes an outer membrane. In some embodiments, the outer membrane is a flow-limiting membrane. In some embodiments, the outer membrane is an analyte-permeable membrane. Petition 870250049834, dated 06 / 13 / 2025, p. 13 / 97 5 / 76

[0022] In some embodiments of the present invention, the measurement of the accumulated charge signal includes measuring a peak signal height and / or measuring a peak signal area.

[0023] In some embodiments, the method also includes calibration of the measured peak height to obtain an analyte concentration.

[0024] In some embodiments, the method also includes calibration of the measured peak area to obtain an analyte concentration.

[0025] In some embodiments, the measurement of the accumulated charge signal comprises 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] In some embodiments of the present invention, the operating electrode includes a detection element comprising the analyte-specific enzyme and the redox mediator. In some embodiments, the detection element also includes carbon nanotubes.

[0027] In some embodiments, a method for the detection of an analyte that makes use of a sensor, wherein the sensor includes an operating electrode that includes an analyte-specific enzyme and a redox mediator, includes: the provision of the operating electrode to the analyte; the accumulation of charge derived from the analyte that reacts with the analyte-specific enzyme and the redox mediator; and the measurement of a signal of the accumulated charge by measuring a peak height of the signal and / or measuring a peak area of ​​the signal.

[0028] In some embodiments of the present invention, a system for the detection of an analyte includes an operating electrode, a detection element disposed in the operating electrode, wherein the detection element includes an analyte-specific enzyme and a redox mediator, and the detection element is configured to accumulate charge derived from the analyte reacting with the analyte-specific enzyme for a stipulated period of time, and a circuit configured to connect Petition 870250049834, dated 06 / 13 / 2025, p. 14 / 97 6 / 76 with the operating electrode after the stipulated time period, and to measure the accumulated charge signal. In some embodiments, the detection element of this system includes carbon nanotubes. In some embodiments, this system also includes an outer membrane that covers at least the detection element. In some embodiments, the analyte-specific enzyme of this system is selected from a nicotinamide adenine dinucleotide (NAD)-dependent dehydrogenase, a flavin adenine dinucleotide (FAD)-dependent oxidase, or a flavin mononucleotide (FMN)-dependent oxidase. For example, in some embodiments, the analyte-specific enzyme of this system is selected from 11β-hydroxy steroid dehydrogenase type 2 (11p-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 is a flowchart describing a method for detecting the accumulation mode which includes actions 10, 15, 20, 25 and 30, as indicated, according to embodiments of the present invention.

[0030] Figure 2 shows a schematic diagram of the electrode configurations used for the accumulation detection mode according to embodiments of the present invention wherein, when the circuit is connected as shown in the left panel, the operating electrode is balanced to a potential (voltage) sufficient to drive the analyte redox reaction under steady-state conditions and, when the circuit is disconnected as shown in the right panel, the operating electrode is electrically disconnected from the circuit, allowing electrons from the analyte to be stored in the redox polymer until the operating electrode is reconnected to the circuit and the stored charge can be measured.

[0031] Figure 3A shows the expected sign of the current versus (vs.) the Petition 870250049834, dated 06 / 13 / 2025, p. 15 / 97 7 / 76 time and certain quantitative parameters (accumulation time when the circuit is broken, peak area and peak height, as indicated) of the accumulation detection mode, according to the embodiments of the present invention.

[0032] Figure 3B shows a schematic diagram of the redox reactions that occur during the accumulation detection mode (when the circuit is broken as shown as a broken circuit as indicated) of an oxidizable analyte (analyte A) when using an oxidase enzyme (AOx) co-immobilized with an osmium redox polymer (Os3+), according to embodiments of the present invention.

[0033] Figure 3C shows the current versus time graph obtained for the 2 μM glucose accumulation detection mode (as indicated in white) using an exemplary glucose sensor (at +40 mV as indicated by the hatched lines) and measured for five different accumulation times, according to embodiments of the present invention.

[0034] Figure 3D shows the calibration curves of the amperometry and accumulation mode signals measured by peak height or peak area for the accumulation times shown in Figure 3C, according to the embodiments of the present invention.

[0035] Figure 4A shows a representative graph of current versus time for a calibration experiment using the accumulation detection mode with an exemplary glucose sensor (at +40 mV as indicated by the hatched lines and a second accumulation time of 60 (when the circuit is broken as indicated in white) for each detection, according to the embodiments of the present invention.

[0036] Figure 4B shows a comparison of the calibration curves resulting from the amperometry and accumulation mode signals measured for the detection experiment shown in Figure 4A, according to the modes of Petition 870250049834, dated 06 / 13 / 2025, p. 16 / 97 8 / 76 present invention.

[0037] Figure 5 shows the calibration curves for amperometry mode detection and accumulation (peak height and peak area) at accumulation times of 1 (diamonds), 2 (triangles), 5 (squares) and 10 (circles) minutes as indicated at glucose concentrations of 0, 50, 100, 200 and 500 μM, where each calibration curve represents the average response of four sensors, according to the embodiments of the present invention.

[0038] Figure 6A shows a graph of the potential signal versus time of a model glucose sensor obtained by using the open-circuit potential method to detect various nanomolar (nM) glucose concentrations as indicated, according to embodiments of the present invention.

[0039] Figure 6B shows a calibration curve (slope versus glucose concentration (nM)) of the data from the graph in Figure 6A, according to embodiments of the present invention.

[0040] Figure 6C shows a graph of the potential signal versus the time signal of a model glucose sensor obtained by using the open-circuit potential method to detect various nM glucose concentrations as indicated, according to embodiments of the present invention.

[0041] Figure 6D shows a calibration curve (slope versus glucose concentration (nM)) of the data from the graph in Figure 6C, according to embodiments of the present invention.

[0042] Figure 6E shows a composite calibration curve for the model glucose sensors (circle data points, n = 8) and control sensors (open circle data points, n = 4) of in vitro glucose detection when using the open circuit potential method, according to embodiments of the present invention.

[0043] Figure 6F shows an approximation of the calibration curve of Petition 870250049834, dated 06 / 13 / 2025, page 17 / 97 9 / 76 Figure 6E shows glucose concentrations from 0 to 200 nM, according to embodiments of the present invention.

[0044] Figure 6G shows a graph of the potential signal versus time of a model glucose sensor obtained by using the open-circuit potential method with a model glucose sensor as the operating electrode and a control sensor (which has a redox polymer but no glucose oxidase) as the reference electrode, wherein the model glucose sensor is to detect various nM glucose concentrations as indicated, according to embodiments of the present invention.

[0045] Figure 6H shows a calibration curve (slope versus glucose concentration (nM)) of the data from the graph in Figure 6G, according to embodiments of the present invention.

[0046] Figure 7 shows a comparison of the accumulation mode signal shape under different filtering frequencies with 3.2 Hz shown with a solid black line and 0.032 Hz shown with a dashed line, according to embodiments of the present invention.

[0047] Figure 8A shows two micrographs of the glucose detection reagent deposited with carbon nanotubes (right panel) and without carbon nanotubes (left panel), CNTs, according to embodiments of the present invention.

[0048] Figure 8B shows calibration curves for amperometry and accumulation mode detection (peak height and peak area) when using different filtering frequencies (0.032 Hz shown as circles and 3.2 Hz as triangles) and detection reagent with and without CNTs, according to embodiments of the present invention.

[0049] Figure 9A shows the accumulation mode signals obtained for a representative glucose sensor during a calibration experiment using Petition 870250049834, dated 06 / 13 / 2025, page 18 / 97 10 / 76 glucose concentrations from 0 to 200 nM, with an accumulation time of 30 minutes for each detection, a signal filtered at 3.2 Hz, and CNTs added to the detection reagent, according to the embodiments of the present invention.

[0050] Figure 9B shows the calibration curves with the corresponding linear fit resulting from the amperometry and accumulation mode signals measured for the detection experiment shown in Figure 8A, where each signal is the average subtracted from the background of 8 sensors, wherein the error bars represent the standard deviation, and the bottom row of the graphs is an approximation showing glucose concentrations from 0 to 50 nM, according to embodiments of the present invention.

[0051] Figure 10A shows the accumulation mode signals of a representative glucose sensor under background conditions ([glucose] = 0) in a buffer solution open to the atmosphere (bold line) and purged with oxygen (thin line), according to embodiments of the present invention.

[0052] Figure 10B shows a summary of the amperometry and background accumulation mode signals of the experiment shown in Figure 10A wherein the signals are the average (mean) of 4 sensors, and the oxygen-purged data are shown as solid circles and the atmospheric data are shown as open circles, according to embodiments of the present invention.

[0053] Figure 11 shows the calibration curves obtained for amperometry and accumulation mode detection (peak height and peak area) during a detection experiment with glucose concentrations from 0 to 200 μM, wherein the linear lines are shown as the most adapted linear lines obtained for concentrations from 0 to 200 nM that are predicted for the higher concentrations, and each signal is the average of 8 sensors, according to embodiments of the present invention.

[0054] Figure 12 shows a schematic diagram of a sensor. Petition 870250049834, dated 06 / 13 / 2025, p. 19 / 97 11 / 76 analyte according to the embodiments of the present invention.

[0055] Figure 13 is a cross-sectional view showing a portion of an analyte sensor that is compatible with one or more embodiments of the present invention.

[0056] Figure 14A shows a plan view of an implantable analyte sensor that is compatible with one or more embodiments of the present invention.

[0057] Figure 14B is a cross-sectional view illustrating a portion of any analyte sensor having a membrane that is compatible with one or more embodiments of the present invention.

[0058] Figure 14C shows a close-up view of the detection layer, the operating electrode and the substrate with an overlaid outer membrane, according to embodiments of the present invention.

[0059] Figure 14D is a schematic diagram illustrating a redox reaction of an analyte with an analyte-specific enzyme and a redox mediator at an operating electrode, according to embodiments of the present invention.

[0060] Figure 15 is a block diagram of an embodiment of an analyte monitoring system according to the embodiments of the present invention.

[0061] Figure 16 is a block diagram of an embodiment of a reading device for the analyte monitoring system of Figure 15, according to the embodiments of the present invention.

[0062] Figure 17 is a block diagram of an embodiment of a sensor control device for the analyte monitoring system of Figure 15, according to the embodiments of the present invention. DETAILED DESCRIPTION

[0063] The embodiments of the present invention provide a method of electrochemical measurement by using an electrochemical sensor to measure low Petition 870250049834, dated 06 / 13 / 2025, page 20 / 97 12 / 76 nanomolar concentrations of analyte in vitro and in vivo. Embodiments of the present invention include an electrochemical sensor such as a modified enzymatic biosensor for measuring low nanomolar concentrations of an analyte.

[0064] Where a range of values ​​is employed, it should be understood that each intermediate value, to the tenth power of unity of the lower limit, unless the context clearly dictates otherwise, between the upper and lower limits of that range, is also specifically disclosed. Each smaller range between any stated value or the intermediate value in a stated range and any other stated or intermediate value in that stated range is encompassed within the description. The upper and lower limits of these smaller ranges may be independently included or excluded in the range, and each range where any, neither, or both limits are included in the smaller ranges is also encompassed within the description, subject to any limit specifically excluded in the stated range. Where the stated range includes one or both limits, the ranges that exclude either or both of these included limits are also included in the description.

[0065] As used in this document, the terms substantially, about, and similar terms are used as approximation terms and not as degree terms, and are intended to account for the inherent deviations in measured or calculated values ​​that should be recognized by those normally versed in the state of the art.

[0066] In the description as disclosed in this document, it should be understood that a word appearing in the singular encompasses its counterparts in the plural, and a word appearing in the plural encompasses its counterparts in the singular, unless understood implicitly or explicitly or otherwise indicated. Merely by way of example, a reference to one or the analyte encompasses a single analyte, as well as a combination and / or mixture of two or more different analytes, a reference to one or the concentration value Petition 870250049834, dated 06 / 13 / 2025, p. 21 / 97 13 / 76 encompasses a single concentration value, as well as two or more concentration values, and others, unless understood implicitly or explicitly or otherwise indicated. Furthermore, it should be understood that for any component described herein, any of the possible candidates or alternatives listed for that component may generally be used individually or in combination with one another, unless understood implicitly or explicitly or otherwise indicated. In addition, it should be understood that any list of such candidates or alternatives is merely illustrative and not limiting, unless understood implicitly or explicitly or otherwise indicated.

[0067] As used in this document, the terms measure, measuring and measured may encompass the meaning of a respective of the terms determines determining, determined, calculates, calculating and calculated.

[0068] As used in this document, an electrochemical sensor is a device configured to detect the presence and / or measure the level of an analyte in a sample through electrochemical oxidation-reduction reactions in the sensor. These reactions are transduced into an electrical signal that can be correlated to a quantity, concentration, or level of an analyte in the sample.

[0069] As used in this document, an operating electrode is an electrode in which the analyte (or a second compound whose level depends on the level of the analyte) is electro-oxidized or electro-reduced with or without the interference of an electron transfer agent.

[0070] As used in this document, a counter electrode refers to an electrode paired with the working electrode, through which a current passes that is equal in magnitude and opposite in sign to the current passing through the working electrode. In the context of embodiments of the present invention, the term counter electrode includes both a) counter electrodes and b) counter electrodes that also Petition 870250049834, dated 06 / 13 / 2025, p. 22 / 97 14 / 76 function as reference electrodes (i.e., counter electrodes / reference electrodes), unless otherwise indicated.

[0071] As used in this document, a reference electrode includes both a) reference electrodes and b) reference electrodes that also function as counter electrodes (i.e., counter electrodes / reference electrodes), unless otherwise indicated.

[0072] As used in this document, electrolysis is the electrooxidation or electroreduction of a compound directly at an electrode or through one or more electron transfer agents.

[0073] As used in this document, components are immobilized within a sensor, for example, when the components are retained in or covalently, ionically or coordinately bonded to the sensor constituents and / or are retained in a polymeric or sol-gel matrix or membrane that prevents mobility.

[0074] As used in this document, an electron transfer agent is a compound that carries electrons between the analyte and the operating electrode, either directly or in cooperation with other electron transfer agents. An example of an electron transfer agent is a redox mediator.

[0075] As used in this document, a redox mediator is an electron transfer agent for carrying electrons between an analyte, an analyte-reduced or analyte-oxidized enzyme, and an electrode, either directly or through one or more additional electron transfer agents. A redox mediator that includes a polymeric backbone may also be referred to as a redox polymer.

[0076] As used in this document, the term precursor polymer refers to the initial polymer before various modifying groups. Petition 870250049834, dated 06 / 13 / 2025, p. 23 / 97 15 / 76 are joined together to form a modified polymer.

[0077] As used in this document, a detection layer is a sensor component that includes constituents that facilitate the electrolysis of the analyte. The detection layer may include constituents such as an electron transfer agent (e.g., a redox mediator or a redox polymer), a catalyst (e.g., an analyte-specific enzyme) that catalyzes an analyte reaction to produce a response at the operating electrode, or an electron transfer agent and a catalyst. In some embodiments of the present invention, a sensor includes a detection layer that is arranged in a non-leachable manner in the vicinity of or on the operating electrode.

[0078] As used in this document, a detection element is an application or region of an analyte-specific enzyme arranged with the detection layer. In this way, a detection element is able to interact with the analyte. A detection layer may have more than one detection element composing the analyte detection area arranged on the operating electrode. In some embodiments, the detection element includes an analyte-specific enzyme and an electron transfer agent (e.g., redox mediator). In some embodiments, the detection element includes an analyte-specific enzyme, an electron transfer agent, and a crosslinker.

[0079] As used in this document, a non-leachable or non-releasable compound, or a compound that is disposed of in a non-leachable manner, means a compound that is affixed to the sensor in such a way that it does not diffuse away substantially from the detection layer of the operating electrode for the period during which the sensor is used (for example, the period during which the sensor is implanted in a patient or during which a sample is measured).

[0080] As used in this document, a crosslinker is a molecule that contains at least two reactive groups that can bind at least Petition 870250049834, dated 06 / 13 / 2025, p. 24 / 97 16 / 76 two molecules to each other, or linking at least two portions of the same molecule to each other. The linking of at least two molecules is called intermolecular crosslinking, while the linking of at least two portions of the same molecule is called intramolecular crosslinking. A crosslinker that has more than two reactive groups may have the ability to crosslink both intermolecularly and intramolecularly at the same time.

[0081] A membrane solution is a solution that contains all the components necessary for crosslinking and membrane formation, including a modified polymer containing heterocyclic nitrogen groups, a crosslinker, and a buffering agent or a mixed alcohol-buffering agent solvent.

[0082] As used in this document, a biological fluid or biofluid is any body fluid or body fluid derivative in which the analyte can be measured, for example, blood, interstitial fluid, plasma, dermal fluid, sweat and tears.

[0083] As used in this document, accumulation mode detection refers to the accumulation of electrons produced from the oxidation of an analyte, wherein the oxidation occurs in or on the detection element of an operating electrode that is not connected to a circuit, thereby creating the accumulation of electrons. Accumulation Mode Detection

[0084] With reference to the method flowchart in Figure 1, some embodiments of the present invention include a method for obtaining a signal from an analyte using a sensor, wherein the sensor includes an operating electrode and another electrode (for example, a counter electrode and / or reference electrode) wherein the operating electrode is provided or modified with (10) a catalyst such as an analyte-specific enzyme and an electron transfer agent (for example, a redox mediator). The area of ​​the operating electrode that is modified with the enzyme Petition 870250049834, dated 06 / 13 / 2025, p. 25 / 97 17 / 76 specific analyte and the redox mediator can be indicated as the detection element or detection layer of the operating electrode. As shown in Figure 1, the operating electrode that has been provided (e.g., modified) with an analyte-specific enzyme is provided (15) with the analyte. In the presence of the analyte, the modified operating electrode oxidizes the analyte, and the amount of oxidation is measured as the amount of electron charge produced from the reaction. As long as the operating electrode is not connected to another electrode, the charge from the redox reaction will continue to accumulate (20) on the operating electrode. For analytes at a low concentration in the body (e.g., cortisol), the accumulation of charge (electrons) over a stipulated period of time allows low analyte concentrations to result in an output signal that is easy to measure and quantify compared to other known methods.After a stipulated time period for charge accumulation (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 operating electrode is connected (25) with at least one other electrode such as a counter electrode and / or reference electrode to form a circuit. With the formation of the circuit, the electrons accumulated in the operating electrode are discharged as an electrical signal, the amplitude of which is measured (30) and correlated to the amount of analyte present in the operating electrode. In this way, by following the method according to the embodiments of the present invention as shown in actions 10, 15, 20, 25 and 30 of Figure 1, low concentrations (e.g., nanomolar amounts as low as 4.7 nM) of an analyte can be detected and measured immediately.

[0085] With reference to Figure 2, an example of a three-electrode configuration is shown with an operating electrode 40, a reference electrode 50 and a counter electrode 60 used for the accumulation detection mode according to embodiments of the present invention wherein, when the circuit 70 is connected as shown in the left panel, the operating electrode is balanced at a potential Petition 870250049834, dated 06 / 13 / 2025, page 26 / 97 18 / 76 (voltage) sufficient to drive the analyte redox reaction under steady-state conditions. For example, for the exemplary glucose sensor used in this document, the potential (voltage) sufficient to drive the redox reaction is +40 mV versus Ag / AgCl. When circuit 70 is not connected as shown in the right panel, the working electrode 40 is electrically disconnected from circuit 70, allowing the charge (e.g., electrons) of the analyte to be stored in the redox polymer until the working electrode 40 is reconnected to circuit 70 and the stored charge is measured.

[0086] With reference to Figures 3A and 3B, an example of an electrochemical enzymatic biosensor is shown in a conceptual overview of an accumulation mode. In this example, the detection of the analyte (A) is based on the presence of an oxidoreductase enzyme (AOx) electrically wired to the sensor's operating electrode via a redox polymer. During normal amperometric detection, the electrode is balanced at a potential (voltage) so that the analyte is reacted at a constant rate, which is proportional to the analyte concentration. For an oxidation reaction of the analyte (A to A+), as shown in Figure 3B, electrons will flow from the analyte (A) to the analyte-specific enzyme (AOx) to the redox polymer (e.g., Os3+) to the operating electrode at a constant rate, producing a steady-state current as shown in Figure 3A.If the working electrode is disconnected from the circuit, the flow of electrons from the redox polymer to the working electrode will stop, resulting in no current flow through the circuit. However, the analyte will still undergo enzymatic oxidation, which in turn results in the reduction of the redox polymer (Os3+ to Os2+). This results in an accumulation (shown by the Os2+ cloud) of the reduced form of the redox polymer (Os2+) over time, as electrons (e-) from the analyte are stored in the redox polymer. When the working electrode is reconnected to the circuit so that it is balanced at its original potential (voltage), the accumulation... Petition 870250049834, dated 06 / 13 / 2025, page 27 / 97 19 / 76 The reduced form of the redox polymer will be oxidized, resulting in a large current spike as shown in Figure 3A. The current will then decay back to the original amperometric current when the redox system reaches a steady state once again. This two-step process forms the basis for accumulation mode detection: a first step in which the sensor's operating electrode is disconnected from or not connected to the circuit for a specified period of time (also referred to as accumulation time), allowing the analyte charge to accumulate in the redox polymer, and a second step in which the sensor's operating electrode is connected to the circuit after the accumulation time, allowing the accumulated charge to be discharged and measured as a sharp spike.

[0087] With reference to Figures 3C and 3D, an example of accumulation mode detection was demonstrated using a developed glucose sensor consisting of a glucose-specific detection reagent deposited on a screen-printed carbon electrode. The glucose detection reagent consists of the glucose oxidase enzyme cross-linked to an Os redox polymer. This reagent has already been demonstrated for use in glucose biofuel cells, as well as self-energized and potentiostat-energized continuous glucose sensors. See, for example, 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 full texts of which are incorporated herein by reference.In some embodiments of the present invention, a method for detecting accumulation mode can be used to increase the sensitivity of an electrochemical measurement. For the experiment shown in Figures 3C and 3D, a glucose sensor was placed in a solution of 2 μM glucose and 100 mM phosphate-buffered saline (PBS), and several measurements of the accumulation mode were taken. Petition 870250049834, dated 06 / 13 / 2025, page 28 / 97 20 / 76 measurements were taken while the sensor current was monitored. For each measurement, the sensor was initially balanced to +40 mV to drive steady-state oxidation of glucose, the operating electrode was then electrically disconnected for a stipulated period of time (the accumulation time) to allow charge accumulation, and the operating electrode was then reconnected to measure the accumulated charge. As shown, the size of the oxidizing current peak increases with increasing accumulation time. Therefore, by simply increasing the accumulation time (e.g., up to 30 seconds, 60 seconds, or up to 120 seconds), the sensitivity of the measurement with this glucose sensor and the glucose concentration are increased. The amperometric signal, which was measured as steady-state sensor current, as well as the peak height and peak area of ​​the current peaks measured in Figure 3C, are plotted against the accumulation time in Figure 3D.As shown, the amperometric current is not dependent on the accumulation time and does not remain constant. However, the height and area of ​​the current peak show a linear dependence on the accumulation time, highlighting the advantage that accumulation mode detection has over traditional amperometry. That is, the sensor sensitivity can be adjusted by changing an easily adjustable parameter of the measurement technique, for example, the time period for the accumulation charge.

[0088] According to embodiments of the present invention, the accumulation mode detection 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 for glucose concentrations up to 100 μM. As indicated, a time of 60 seconds was used for each detection. Figure 4A shows the resulting current trace over time in this experiment. As shown, both the steady-state amperometric current and the size of the accumulation mode current peaks increase with a Petition 870250049834, dated 06 / 13 / 2025, p. 29 / 97 21 / 76 increasing glucose concentration. Figure 4B shows graphs of amperometric current and peak height and peak area of ​​the current peaks as a function of glucose concentration, with all three signals exhibiting a linear dependence on analyte concentration. Therefore, the results show that the accumulation detection mode, if measured using peak height or peak area, results in linear calibration curves and can thus be used for detection in a manner analogous to traditional amperometry with greater sensitivity. In this way, since the peak height obtained from the accumulation detection mode is measured in current units, the sensitivity of this measurement method can be quantitatively compared to the sensitivity of amperometry. For example, the sensitivity of the measurement method can be obtained by comparing the slopes of the calibration curves, such as those shown in Figure 4B.By comparison, amperometry has a sensitivity of 0.44 nA / μM, whereas the accumulation detection mode (when using peak height measurement) has a sensitivity of 1.69 nA / μM. Therefore, with an accumulation time of 60 seconds, the accumulation detection mode according to the embodiments of the present invention increases the sensitivity of the electrochemical measurement by a factor of about 4 compared to amperometry.

[0089] Furthermore, since both peak height and peak area provide the same result and sensitivity, in some embodiments of the present invention, a means of measuring the current signal resulting from the operating electrode includes calculating the peak height and / or peak area.

[0090] In some embodiments of the present invention, the detection of the accumulation mode is performed using a sensor that has an outer membrane. Since electrochemical sensors are frequently coated with an outer membrane (e.g., a polymer membrane) in order to provide stability to the detection reagents, bulk transport limitations, Petition 870250049834, dated 06 / 13 / 2025, p. 30 / 97 22 / 76 biocompatibility and / or to prevent electrodes from becoming soiled, a polymer-coated sensor was tested to ensure that accumulation mode detection is performed as expected. With reference to Figure 5, an exemplary glucose sensor coated with an outer flow-limiting polymer membrane was used to obtain calibration curves through amperometry and accumulation mode detection at glucose concentrations of 0, 50, 100, 200, and 500 μM. Four consecutive measurements were made at each glucose concentration using a different accumulation time of 1, 2, 5, and 10 minutes as indicated by the data points, respectively, in Figure 5.

[0091] As shown in Figure 5, measurements using amperometry mode (left graph) and accumulation mode (middle and right graphs) provide a linear response to analyte concentration. As expected, when using amperometry (left graph in Figure 5), the sensor sensitivity is independent of the accumulation time. However, when using accumulation mode detection (middle and right graphs in Figure 5), the sensor sensitivity increases with increasing accumulation time. Due to the flow-limiting outer membrane, the sensor sensitivities when using amperometry and accumulation mode detection are much lower than for sensors without an outer membrane. This is expected, since the outer membrane limits the diffusion of the analyte to the detection reagent.However, as shown in Figure 5, the accumulation mode detection is performed as expected when an external polymer membrane is added to the sensor and provides another example of how the sensor's sensitivity can be adjusted by changing the accumulation time. Furthermore, it should be noted that a stipulated time period of more than 10 minutes for charge accumulation when using the accumulation detection mode with continuous monitoring sensors can cause negative effects on the sensor's time setting. Therefore, in some embodiments of... Petition 870250049834, dated 06 / 13 / 2025, p. 31 / 97 23 / 76 of the present invention, the detection of the accumulation mode is carried out by using a sensor that has an outer membrane where the stipulated time period for charge accumulation is up to 10 minutes.

[0092] It should also be noted that, when an outer membrane such as a flow-limiting outer membrane may not be necessary to prevent electrode fouling when measuring analytes at low concentrations, an outer membrane can provide a biocompatible interface with an in vivo environment and / or confer stability to the underlying detection layer including electron transfer agents and / or analyte-specific enzymes therein. For the accumulation detection mode in which an outer membrane is used, the time period stipulated for charge accumulation can be increased to allow for oxidation of the total analyte concentration.In some embodiments of the present invention, a method for detecting the accumulation mode when using a sensor that has an outer membrane includes increasing the stipulated time period for accumulating the charge up to 1 minute, up to 2 minutes, up to 3 minutes, up to 4 minutes, up to 5 minutes, up to 6 minutes, up to 7 minutes, up to 8 minutes, up to 9 minutes, or up to 10 minutes in order to allow the complete reaction of all the analyte present in the operating electrode. In some embodiments of the present invention, a method for detecting the accumulation mode when using a sensor that has an outer membrane includes increasing the stipulated time period for accumulating the charge from 10 minutes up to 30 minutes.

[0093] Alternatively, in some embodiments of the present invention, the outer membrane may be made of a highly permeable material and, thus, although the permeable membrane does not attenuate the rate at which the analyte reaches the detection layer of the operating electrode, the permeable membrane allows for stability, bulk transport limitations, and / or biocompatibility. Non-limiting examples of highly permeable membrane materials include Petition 870250049834, dated 06 / 13 / 2025, page 32 / 97 24 / 76 poly(vinylpyridine) crosslinked with high molecular weight (MW > 400 g / mol) poly(ethylene glycol) diglycidyl ether, derivatized poly(vinylpyridine) crosslinked with high molecular weight (MW > 400 g / mol) poly(ethylene glycol) diglycidyl ether), poly(vinyl) alcohol, poly(acrylic) acid, and poly(methacrylic) acid.

[0094] With reference to Figures 6A-6B, an electrochemical glucose sensor was used in the in vitro experiment to measure (i.e., detect) glucose concentrations ranging from 0 to 1,000 nanomolar (nM) glucose. In this example, the operating electrode of the sensor included the glucose oxidase enzyme crosslinked to an Os-based redox polymer deposited and immobilized on a screen-printed carbon electrode. The experiment was performed as disclosed in this document (e.g., in Example 8). In addition, a screen-printed carbon counter electrode and an Ag / AgCl reference electrode were used. Before each measurement, the operating electrode was held at +40 mV versus (vs.) Ag / AgCl for 3 minutes, after which the electrode open-circuit potential was measured for 3 minutes. The graph in Figure 6A shows the resulting potential versus time curve for the indicated glucose concentrations (0 to 1,000 nM glucose).Therefore, as shown, higher glucose concentrations result in a rate of change of potential of greater magnitude. In some embodiments of the present invention, the rate of change is calculated as the slope of the potential versus time curve. Figure 6B is a calibration curve showing a graph of the rate of change (calculated as the slope from 30 to 180 seconds) versus glucose concentration. As shown in Figure 6B, the rate of change of potential shows a linear dependence on glucose concentration.

[0095] With reference to Figures 6C-6C, the same electrochemical glucose sensor used in the experiment of Figures 6A-6B was used in the in vitro experiment to measure glucose concentrations ranging from 0 to 750 nM of glucose, including glucose concentrations below 100 nM (e.g., 10 nM, 25 nM, and Petition 870250049834, dated 06 / 13 / 2025, page 33 / 97 25 / 76 nM). The graph in Figure 6C shows the resulting potential versus time curve for the indicated glucose concentrations. Therefore, as shown in Figure 6D, the rate of change plotted for this experiment remains linear up to 10 nM of glucose. This correlation is also shown in Figure 6E, which shows a calibration curve resulting from testing 8 individual glucose sensors. Furthermore, control sensors lacking the glucose oxidase enzyme (but still possessing the Os redox polymer) were also tested in this experiment. As shown in Figures 6E and 6F, the rate of change of the control sensors represented by the open circles showed no dependence on glucose concentrations.

[0096] According to some embodiments of the present invention, the method disclosed herein can be used to lower the background signal (e.g., signal at [analyte] = 0). With reference to Figures 6G-6H, an experiment was performed using the glucose sensor used in the experiment shown in Figure 6A as the operating electrode. In addition, a control sensor that does not have the glucose oxidase enzyme, but still has the Os redox polymer, was used as the reference electrode during the open circuit potential measurement. By using this configuration, the amount of measured signal that is not from glucose oxidation is minimized. For example, when using a control sensor without glucose oxidase as the reference electrode, the background signal (the slope curve of the potential versus time for a glucose concentration equal to zero) is approximately equal to zero.The resulting 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, which was obtained using an Ag / AgCl reference electrode. Therefore, the methods and systems of the present invention include the use of a glucose oxidase-free control sensor as a reference electrode during open circuit potential measurement as an effective method for reducing the signal. Petition 870250049834, dated 06 / 13 / 2025, page 34 / 97 26 / 76 background.

[0097] In some embodiments of the present invention, a signal produced from the redox reaction of an analyte in the detection layer of an operating electrode can be adjusted or modified to enhance the output signal for any concentration of the sensor and / or analyte. In some embodiments of the present invention, the signal is enhanced by modifying the frequency at which the current signal is recorded. For example, with reference to Figure 7, in order to maximize the peak height measured during the peak current of accumulation detection, the signal can be recorded at a faster sampling rate (e.g., 0.1 Hz) and filtered at a higher frequency (e.g., 3.2 Hz) than the sampling rate of 0.5 Hz and the filter frequency of 0.03 Hz that were used for the accumulation mode detection experiments disclosed herein and shown in Figures 3A to 3D, 4A-4B and 5.As shown in Figure 7, the detection peak is much sharper at the higher frequency of 3.2 Hz, leading to a higher peak height. Therefore, in some embodiments of the present invention, the method of detecting the accumulation mode includes increasing the filter frequency up to 3.2 Hz to maximize the signal magnitude. It should be noted that at a frequency higher than 3.2 Hz, the signal-to-noise ratio is too large to allow accurate measurements if amperometric current or accumulation peak measurement is used.

[0098] In some embodiments of the present invention, carbon nanotubes (CNTs) are added to the detection element of the operating electrode. For example, CNTs are added to the detection reagent including the redox mediator and the analyte-specific enzyme and applied to the operating electrode. With reference to Figure 8A, CNTs were added to the detection reagent in the micrograph on the right and CNTs were not added in the micrograph on the left. Accumulation mode detection was measured with and without CNTs. As shown Petition 870250049834, dated 06 / 13 / 2025, p. 35 / 97 27 / 76 in Figure 8B, with the addition of CNTs with the detection element on the operating electrode, the peak current of the accumulation mode has a greater peak height.

[0099] In some embodiments of the present invention, the accumulation mode detection includes the use of a sensor with an accumulation time (e.g., a stipulated time period for charge accumulation) of 30 minutes, a signal frequency filter at 3.2 Hz, and the addition of carbon nanotubes (CNTs) to the detection element on the operating electrode. Figure 9A shows the accumulation mode signals obtained for a representative glucose sensor at glucose concentrations from 0 to 200 nM in the presence of CNTs, with an accumulation time of 30 minutes, and the signal filtered at 3.2 Hz. Therefore, as shown in the signal calibration curves in Figure 9B, compared to amperometry, the accumulation mode detection according to the embodiments of the present invention provides increased sensitivity for analytes at low concentrations.As seen, with an accumulation time of 30 minutes, the accumulation detection mode using peak height measurement provides an 800-fold increase in sensitivity compared to amperometry. Regarding the detection limit, the accumulation detection mode using peak area measurement is superior, resulting in a lower detection limit (LOD) of 4.7 ± 1.4 nM, a 25-fold increase compared to amperometry. Although the linear range for accumulation mode detection is more limited than for amperometry, it should be noted that this range can be shifted to higher concentrations by using a shorter accumulation time. Sensor for detecting accumulation mode.

[0100] A sensor as described herein may be an in vivo sensor or an in vitro sensor (i.e., a distinct monitoring test strip). Such a sensor may be formed on a substrate, for example, a substantially planar substrate. In certain embodiments, the sensor is a wire, for example. Petition 870250049834, dated 06 / 13 / 2025, p. 36 / 97 28 / 76 example, an inner portion of the operating electrode wire with one or more other electrodes associated with it (e.g., on, including wrapped around). The sensor may also include at least one counter electrode (or counter electrode / reference electrode) and / or at least one reference electrode or at least one reference electrode / counter electrode.

[0101] Figure 12 schematically illustrates an embodiment of an analyte sensor 800 according to the embodiments of the present invention. This sensor includes electrodes 801, 802, and 803 on a base 804. The electrodes (and / or other elements) can be applied or processed using any appropriate technology, for example, chemical vapor deposition (CVD), physical vapor deposition, ion bombardment, reactive ion bombardment, printing, coating, ablation (e.g., laser ablation), painting, dip coating, etching, and others.The 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 an amalgam), 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 metallic compounds of these elements.

[0102] The analyte sensor 800 may be fully implantable in a user, or it may be configured so that only one portion is positioned inside (internally) a user and another portion outside (externally) a user. For example, the sensor 800 may include a first portion positionable above a skin surface 810, and a second portion positioned below the skin surface. In such embodiments, the external portion may include contacts (connected to the respective electrodes of the second portion by traces) to connect to another device also external to the user, such as a unit. Petition 870250049834, dated 06 / 13 / 2025, page 37 / 97 29 / 76 transmitter. Although the configuration in Figure 12 shows three electrodes 801, 802, and 803 side by side on the same surface of the base 804, other configurations are contemplated, for example, fewer or more electrodes, some or all electrodes on different surfaces of the base or present on another base, with some or all electrodes stacked together, electrodes of different materials and dimensions, etc.

[0103] Figure 13 shows a cross-sectional view of an embodiment of an analyte sensor 500 having a first portion (which in this embodiment can be characterized as a main portion) positionable above a skin surface, and a second portion (which in this embodiment can be characterized as a minor portion) that includes a sensor tail 530 (which can also be indicated in this document as an insertion tip) positionable below the skin surface (e.g., penetrating through the skin (dermis) and into the subcutaneous space and in contact with the user's biofluid, such as interstitial fluid). The electrode contacts (not shown) are positioned on the first portion of the sensor 500 located above the skin surface and extend to a location on the sensor tail 530.An operating electrode 501, a reference electrode 502, and a counter electrode 503 are shown in the second portion of the sensor 500, and in particular in the lower portion of the sensor tail 530. It should be understood that more or fewer electrodes may be provided in a sensor without departing from the scope of the present invention. For example, a sensor may include more than one operating electrode and / or the counter electrodes and reference electrodes may be a single counter electrode / reference electrode, and so on.

[0104] With regard to Figure 13, the sensor 500 includes a substrate (or substrate layer) 504 and a first conductive layer 508, such as carbon, gold, etc., which is in electrical communication with the detection area 509, thus collectively defining the operating electrode 501. The detection area 509 can be Petition 870250049834, dated 06 / 13 / 2025, p. 38 / 97 30 / 76 protected from microorganisms by providing in one or more components of sensor 500 a microbicidal quality, intended to protect the user's skin health and / or to protect the detection area 509 from potential interference with such microorganisms (e.g., formation of a biofilm due to potential migration of microorganisms). The various electrodes and detection areas defined in the lower portion of the sensor tail 530 in Figure 13 may collectively comprise a detection region, and a microbicidal quality conferred to the sensor tail described herein is provided in the upper portion (upper 25%) of the sensor tail 530 above said region (e.g., above detection area 509, or above electrode 503).

[0105] A first insulating layer 505, such as a first dielectric layer in some embodiments, may be arranged or stacked on at least a portion of the first conductive layer 508, and furthermore a second conductive layer 511 may be arranged or stacked on top of at least a portion of the first insulating layer (or dielectric layer) 505. As shown in Figure 13, the second conductive layer 511 together with a second conductive material 510, such as a silver / silver chloride (Ag / AgCl) layer, may provide the reference electrode 502. Another possible arrangement of the second conductive material 510 is shown in Figure 14B, together with an outer membrane 520 that overlaps the various layers.

[0106] A second insulation layer 506, such as a second dielectric layer in some embodiments, may be arranged or stacked over at least a portion of the second conductive layer 511. In addition, a third conductive layer 513 may be arranged over at least a portion of the second insulation layer 506 and may provide the counter electrode 503. Finally, a third insulation layer 507 may be arranged or stacked over at least a portion of the third conductive layer 513. In this way, the sensor 500 may be Petition 870250049834, dated 06 / 13 / 2025, page 39 / 97 31 / 76 stacked in such a way that a portion of each of the conductive layers is separated by at least one respective insulating layer (e.g., a dielectric layer). Another possible layer configuration is shown in Figure 14B. The embodiments in Figures 13 and 14B show layers having different lengths; however, some or all layers may have the same length and / or width or different lengths and / or widths, without departing from the scope of the present invention.

[0107] In any or all embodiments, some or all of electrodes 501, 502, and 503 may be provided on the same side of substrate 504 in the layered construction described above or, alternatively, may be provided in a coplanar manner such that two or more electrodes may be positioned in the same plane (e.g., side by side, parallel, or angled relative to each other) on substrate 504. For example, coplanar electrodes may include appropriate spacing between them and / or include a dielectric material or insulating material disposed between the conductive layers / electrodes. Furthermore, in some embodiments, one or more of electrodes 501, 502, and 503 may be arranged on opposite sides of substrate 504. In such embodiments, contact pads may be on the same side or on different sides of the substrate.For example, an electrode might be on one side and its corresponding contact might be on the other side; for instance, a trace connecting the electrode and the contact could cross through the substrate.

[0108] With reference now to Figure 14A, another embodiment of an analyte sensor according to one or more embodiments of the present invention is shown, and represents a variation of the sensor 500 of Figures 13 and 14B. With respect to Figure 14A, an implantable detection region (e.g., subcutaneous or transcutaneous) 920 is shown according to one or more embodiments of the present invention including an operating electrode 922 with detection elements. Petition 870250049834, dated 06 / 13 / 2025, p. 40 / 97 32 / 76 931. The proximal end 940 is configured to be connected to various electrical connections to transmit the output signals from the sensing region 920. Collectively, the distal end 925 and the proximal end 940 form the sensor tail. The sensing region 920 encompasses a lower portion of the sensor tail. As described, the sensing region 920 comprises a rounded tip, but other tip shapes may alternatively be present to facilitate insertion into a user's skin.

[0109] In addition, in one or more embodiments, the detection region 920 may include a reference electrode, a counter electrode, or counter-electrodes, such as those shown in Figures 13 and 14B. Alternative electrode configurations may be employed without departing from the scope of the present invention.

[0110] With reference to Figures 13, 14A and 14B, it can be observed that the sensor (or detection region) 500, 920 includes detection functionality in a distal portion of their respective sensor tails. As described above, this location can allow enhanced contact with deeper locations below a user's skin (e.g., the subcutaneous space), where greater access to the user's interstitial fluid can allow greater access to the analyte of interest being measured (e.g., its concentration). That is, the detection region is placed within a user's skin at a sufficient depth to allow accurate measurement of the particular analyte, whereas placing the detection region closer to the skin surface might be inadequate for accurately determining the concentration or other characteristic of a desired analyte.

[0111] With reference to Figures 13 and 14B to 14D, one or more embodiments of the present invention include an operating electrode 501 or 320 having a detection area 509, wherein the detection area 509 has at least one Petition 870250049834, dated 06 / 13 / 2025, p. 41 / 97 33 / 76 detection element 322 which 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 which is positioned in contact with and between 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 disposed 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 disposed in contact with a surface of the reference electrode 502 that is not in contact with the first layer. insulator 505.

[0112] Figure 14C also shows a detection element 322, arranged at least in a portion of the operating electrode 320. In some embodiments of the invention, two or more detection elements 322 may be provided in a detection layer of the operating electrode, where the two or more detection elements are arranged laterally to each other.

[0113] In some embodiments of the present invention, any suitable configuration of the sensing elements 322 may be arranged on the operating electrode 320. Additional configurations of sensing elements are disclosed, for example, in Hoss et al., (US 2012 / 0150005), the full content of which is incorporated herein by reference.

[0114] In some embodiments of the present invention, with reference to Figure 14B, a sensor 500 includes an outer membrane 520 that overlaps at least one operating electrode 501 and the detection area 509. In other embodiments, the outer membrane 520 covers the entire sensor 500. In some embodiments, the outer membrane 520 overlaps all active areas of the sensor 500. For example, the active areas of the sensor 500 are located in the detection region 920 such that Petition 870250049834, dated 06 / 13 / 2025, p. 42 / 97 34 / 76 as shown in Figure 14A and in the detection area 509 as shown in Figure 14B. In some embodiments, the outer membrane 520 overlaps the operating electrode, counter electrode and / or reference electrode in the detection region 920 or in the detection area 509.

[0115] Figure 14C shows a close-up perspective of an outer membrane 335 that overlaps the sensing element 322 disposed on an operating electrode 320 which is disposed on a substrate 325. As shown, the outer membrane 335 is in the process of being overlapped. The outer membrane 335 overlaps at least the entire sensing element 322. Analyte-Specific Enzymes and Electron Transfer Agent (Redox Mediator)

[0116] In some embodiments of the present invention, the sensors of the present invention cannot measure the analyte directly. That is, the electrodes in the sensor cannot interact directly with the analyte. Therefore, the analyte is detected by an enzyme protein that can interact directly with the analyte molecule. However, some enzymes (e.g., glucose oxidase) cannot exchange electrons directly with the electrodes because their redox active sites are buried deep within the enzyme protein structure. Therefore, in order to transfer electrons between the enzyme's redox active site and the electrodes, an electron transfer agent (i.e., a redox mediator) is used. Immobilization of the electron transfer agent and the analyte-specific enzyme in the detection layer creates what is known as a wire since the immobilized molecules are able to relay electrons, and in this way they are electrically wired.An analyte-specific enzyme is also known as a spun enzyme. Spun enzymes are disclosed, for example, in Gregg et al., (US Patent No. 5,262,035), Say et al., (US Patent No. 6,134,461), and Hoss et al., (US Patent Publication No. 2012 / 0150005), the full contents of which are... Petition 870250049834, dated 06 / 13 / 2025, p. 43 / 97 35 / 76 incorporated into this document by way of reference. In some embodiments, the analyte-specific enzyme is cross-linked to the electron transfer agent.

[0117] In some embodiments of the present invention, the electron transfer agents (e.g., redox mediators) are electroreducible and electrooxidizable ions or molecules that have redox potentials (voltage) that are a few hundred millivolts above or below the redox potential (voltage) of the standard calomel electrode (SCE). In some embodiments, the electron transfer agents reduce by no more than about -150 mV and oxidize by no more than about +400 mV versus the SCE. Examples of suitable redox mediators in the form of redox polymers are disclosed, for example, in Mao et al. (US Patent No. 6,605,200), the entire contents of which are incorporated herein by reference.

[0118] According to embodiments of the present invention, with reference to Figure 14D, an electron transfer agent 324 is immobilized on the working electrode 320. In some embodiments, the electron transfer agent 324 and an analyte-specific enzyme 323 are both immobilized on the working electrode 320 by any appropriate means. In some embodiments, the electron transfer agent and the analyte-specific enzyme are co-immobilized on the working electrode with any appropriate crosslinker. In some embodiments, the electron transfer agent and the analyte-specific enzyme are co-immobilized with a chemical crosslinker, for example, (poly(ethylene glycol) diglycidyl ether (PEGDGE)).

[0119] In some embodiments of the present invention, an electron transfer agent to be used in the detection of the accumulation mode includes a redox species selected from osmium, ruthenium, iron, or cobalt coupled with a polymer selected from poly(vinylpyridine), poly(thiophene), poly(aniline), poly(pyrrole), or poly(acetylene). In some embodiments, an electron transfer agent Petition 870250049834, dated 06 / 13 / 2025, page 44 / 97 36 / 76 is a poly(vinylpyridine) osmium (Os) redox polymer of formula I. Formula 1

[0120] In some embodiments of the present invention, the electron transfer agent may be organic, organometallic, or inorganic. Examples of organic redox species are quinones and species that in their oxidized state have quinoid structures, such as Nile blue and indophenol. Some partially oxidized quinones and quinhydrones react with protein functional groups such as the thiol groups of cysteine, the amine groups of lysine and arginine, and the phenolic groups of tyrosine, which may render these redox species unsuitable for some of the sensors of the present invention because of the presence of interfering proteins in an analyte-containing fluid. It should be noted that most substituted quinones and molecules with quinoid structures are less reactive with proteins. In some embodiments, a tetrasubstituted quinone has carbon atoms in positions 1, 2, 3, and 4. Petition 870250049834, dated 06 / 13 / 2025, p. 45 / 97 37 / 76

[0121] Electron transfer agents suitable for use in an accumulation mode detection method according to embodiments of the invention have structures or charges that impede or substantially reduce the diffusional loss of the electron transfer agent during the time period in which the sample is being analyzed. In some embodiments of the present invention, an electron transfer agent includes a redox species linked to a polymer that can be immobilized in the detection layer of the operating electrode. The linkage between the redox species and the polymer can be covalent, coordination, or ionic. Useful electron transfer agents and methods for producing them are described in U.S. Patents Nos. 5,264,104; 5,356,786; 5,262,035; and 5,320,725, the full contents of which are incorporated herein by reference.Although any organic or organometallic redox species can be linked to a polymer and used as an electron transfer agent, in some embodiments of the present invention, the redox mediator is a transition metal compound or complex. In some embodiments, the transition metal compounds or complexes include osmium, ruthenium, iron, and cobalt compounds or complexes. It should be recognized that many of the redox mediator species described herein can also be used, for example, without a polymeric component, as electron transfer agents in a carrier fluid or in a sensor detection layer where leaching of the electron transfer agent is acceptable.

[0122] One type of non-releasable polymeric electron transfer agent contains a redox species covalently bonded within a polymeric composition. An example of this type of mediator is poly(vinyl ferrocene).

[0123] Another type of non-releasable electron transfer agent contains an ionically linked redox species. Typically, this type of mediator includes a charged polymer coupled to a redox species of opposite charge. The Petition 870250049834, dated 06 / 13 / 2025, p. 46 / 97 38 / 76 Examples of this type of mediator include a negatively charged polymer such as Nafion (Dupont) coupled to a positively charged redox species such as a polypyridyl cation coupled with osmium, ruthenium, iron, or cobalt. Another example of an ionically linked mediator is a positively charged polymer, such as quaternized poly(4-vinylpyridine) or poly(1-vinylimidazole), coupled to a negatively charged redox species such as ferricyanide or ferrocyanide. In some embodiments of the present invention, a linked redox species is a highly charged redox species bound within an oppositely charged redox polymer.

[0124] In another embodiment of the invention, suitable non-releasable electron transfer agents include a redox species coordinately linked to a polymer. For example, the mediator may be formed by the coordination of an osmium or cobalt 2,2'-bipyridyl complex to poly(1-vinylimidazole) or poly(4-vinylpyridine).

[0125] In some embodiments of the present invention, the electron transfer agents are osmium transition metal complexes with one or more ligands, wherein each ligand has a nitrogen-containing heterocycle, such as 2,2'-bipyridine, 1,10-phenanthroline, or derivatives thereof. Furthermore, in some embodiments, the electron transfer agents have one or more ligands covalently linked in a polymer, wherein each ligand has at least one nitrogen-containing heterocycle, such as pyridine, imidazole, or derivatives thereof. These preferred electron transfer agents exchange electrons rapidly with each other and with the working electrode so that the complex can be rapidly oxidized and reduced.

[0126] In some embodiments of the present invention, an electron transfer agent includes (a) a polymer or copolymer having pyridine or imidazole functional groups and (b) osmium cations complexed with two Petition 870250049834, dated 06 / 13 / 2025, p. 47 / 97 39 / 76 ligands, wherein each ligand contains 2,2'-bipyridine, 1,10-phenanthroline, or derivatives thereof, and the two ligands are not necessarily the same. In some embodiments, the 2,2'-bipyridine derivatives used for complexation with the osmium cation are 4,4'-dimethyl-2,2'-bipyridine and mono-, di-, and polyalkoxy-2,2'-bipyridines, such as 4,4'-dimethoxy-2,2'-bipyridine. In some embodiments, the 1,10-phenanthroline derivatives used for complexation with the osmium cation are 4,7-dimethyl-1,10-phenanthroline and mono-, di-, and polyalkoxy-1,10-phenanthrolines, such as 4,7-dimethoxy-1,10-phenanthroline. In some embodiments of the present invention, the polymers for complexation with the osmium cation include poly(1-vinylimidazole) polymers and copolymers (referred to as PVI) and poly(4-vinylpyridine) polymers and copolymers (referred to as PVP). Suitable poly(1-vinylimidazole) copolymer substituents include acrylonitrile, acrylamide, and substituted or quaternized N-vinylimidazole.In some embodiments, the electron transfer agents include osmium complexed to a polymer or a copolymer of poly(1-vinylimidazole).

[0127] According to embodiments of the present invention, the electron transfer agents have a redox potential (voltage) ranging from -100 mV to about +150 mV versus the standard calomel electrode (SCE). More specifically, the potential (voltage) of the electron transfer agent ranges from -100 mV to +150 mV. In some embodiments, the potential (voltage) ranges from 50 mV to +50 mV. In other embodiments of the present invention, the electron transfer agents have osmium, ruthenium, iron, or cobalt redox centers and a redox potential (voltage) ranging from +50 mV to -150 mV versus the SCE. Examples of Analyte-Specific Enzymes

[0128] In some embodiments of the present invention, an analyte-specific enzyme is provided (e.g., immobilized) on the operating electrode in order to catalyze the oxidation of the analyte to be measured. As used herein, an analyte-specific enzyme may also be indicated as a Petition 870250049834, dated 06 / 13 / 2025, p. 48 / 97 40 / 76 Analyte Oxidation Enzyme. In some embodiments of the present invention, the analyte-specific enzyme is selected from glucose oxidase, NAD-glucose dehydrogenase, and FAD-glucose dehydrogenase for glucose oxidation. In some embodiments, the analyte-specific enzyme is lactate oxidase or NAD-lactate dehydrogenase for lactate oxidation. In some embodiments, the analyte-specific enzyme is NAD-3-hydroxybutyrate dehydrogenase for 3-hydroxybutyrate oxidation. In some embodiments, the analyte-specific enzyme is 11β-hydroxysteroid dehydrogenase type 2 for cortisol oxidation. In some embodiments, the analyte-specific enzyme is N-alcohol dehydrogenase for alcohol oxidation. In some embodiments, the analyte-specific enzyme is pyruvate oxidase for pyruvate oxidation. In some embodiments, the analyte-specific enzyme is NAD-glutamate dehydrogenase for glutamate oxidation.In some embodiments, the analyte-specific enzyme is xanthine oxidase for theophylline oxidation.

[0129] As should be understood by someone normally versed in the state of the art, any nicotinamide adenine dinucleotide (NAD) or flavin oxidase enzyme can be coupled or immobilized to the detection layer of the operating electrode in order to oxidize its corresponding analyte substrate.

[0130] In some embodiments of the present invention, 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-dicarboxylate dehydrogenase, (R,R)butanediol dehydrogenase, (R)-2-hydroxy fatty acid dehydrogenase, (R)-2-hydroxy acid dehydrogenase, (R)-4-hydroxyphenyl lactate dehydrogenase, (R)-aminopropanol dehydrogenase, (R)-dehydropantoate dehydrogenase, (S,S)-butanediol dehydrogenase, (S)-2-hydroxy fatty acid dehydrogenase, (S)-carnitine 3 Petition 870250049834, dated 06 / 13 / 2025, page 49 / 97 41 / 76 dehydrogenase, (S)-usinate reductase, 1,2-dihydroxy-6-methylcyclohexa-3,5-diene carboxylate dehydrogenase, 1,3-propanediol dehydrogenase, 1,6-dihydroxycyclohexa-2,4-diene-1-carboxylate dehydrogenase, 2-(R)-hydroxypropyl-CoM dehydrogenase, 2-(S)-hydroxypropyl-CoM dehydrogenase, 2-alkenal reductase, 2-alkyn-1-ol dehydrogenase, 2-aminobenzene sulfonate 2,3-dioxygenase, 2-chlorobenzoate 1,2-dioxygenase, 2-coumarate reductase, 2-dehydro-3-deoxy-D-gluconate 2-hydroxy-3-oxopropionate reductase, 2-hydroxybiphenyl-3-monooxygenase, 2-hydroxymethylglutarate dehydrogenase, 2-hydroxyquinoline 5,6-dioxygenase, 2-α-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-coumate dehydrogenase, 2,4-diaminopentanoate dehydrogenase, 2,6-dihydroxypyridine 3-monooxygenase, 2'-phosphotransferase,3-(imidazol-5-yl)lactate dehydrogenase, 3-deamino-3-oxonicotianamine reductase, 3-dehydro-L-gulonate 2-dehydrogenase, 3-hydroxy-2-methylbutyryl-CoA dehydrogenase, 3-hydroxy-2-methylpyridine carboxylate dioxygenase, 3-hydroxyacyl-CoA dehydrogenase, 3-hydroxybenzoate 6-monooxygenase, 3-hydroxybutyrate dehydrogenase, 3-hydroxyisobutyrate dehydrogenase, 3-hydroxyphenyl acetate 6-hydroxylase, 3-hydroxypimeloyl-CoA dehydrogenase, 3-hydroxypropionate dehydrogenase, 3-methylbutanal reductase, a 3-oxoacyl-(acylcarrier-protein) reductase (NADH), 3-phenylpropanoate dioxygenase, 3(or 17)-hydroxysteroid dehydrogenase, 3-alpha-hydroxy-5-beta-androstan-17-one, 3-alpha-hydroxycholanate dehydrogenase, alpha-hydroxysteroid dehydrogenase (specific to A), alpha-hydroxysteroid dehydrogenase (specific to B), 3-alpha,7-alpha,12-alpha-trihydroxycholestan-26-al 26-oxidoreductase, 3-alpha(17-beta)-hydroxysteroid dehydrogenase (NAD+),3-alpha (or 20-beta)-hydroxysteroid dehydrogenase, 3β-hydroxysteroid dehydrogenase, 4-(hydroxymethyl)benzene sulfonate dehydrogenase, 4-aminobenzoate 1-monooxygenase, 4-chlorophenyl acetate 3,4-dioxygenase, 4, Petition 870250049834, dated 06 / 13 / 2025, page 50 / 97 42 / 76 formylbenzene sulfonate dehydrogenase, 4-hydroxytetrahydrodipicolinate reductase, 4-hydroxybenzaldehyde dehydrogenase, 4-hydroxybenzoate 1-hydroxylase, 4-hydroxybenzoate 3-monooxygenase (NAD(P)H), 4-hydroxybutyrate dehydrogenase, 4-hydroxycyclohexanecarboxylate dehydrogenase, 4-hydroxymuconic semialdehyde dehydrogenase, 4-hydroxyphenyl acetaldehyde dehydrogenase, 4-hydroxyphenyl acetate 1-monooxygenase, 4-hydroxyquinoline 3-monooxygenase, 4-hydroxythreonine-4-phosphate dehydrogenase, 4-nitrophenol 2-mono-oxygenase, 4-oxoproline reductase, 4-phosphoerythronate dehydrogenase, 4-sulfobenzoate 3,4-dioxygenase, 4-trimethylammonium butyraldehyde dehydrogenase, 5-carboxymethyl-2-hydroxymuconic-semialdehyde dehydrogenase, 5,6-dihydroxy-3-methyl-2-oxo-1,2,5,6-tetrahydroquinoline dehydrogenase, 6-endohydroxycineole dehydrogenase, 6-hydroxyhexanoate dehydrogenase, 6,7-dihydropteridine reductase, 7-alpha-hydroxysteroid dehydrogenase,15-hydroxyicosatetraenoate dehydrogenase, 15-hydroxyprostaglandin dehydrogenase (NAD+), 15-oxoprostaglandin 13-oxidase, 16-alpha-hydroxysteroid dehydrogenase, 17β-hydroxysteroid dehydrogenase, 20-alpha-hydroxysteroid dehydrogenase, 21-hydroxysteroid dehydrogenase (NAD+), ADP-glyceromannanoheptose 6-epimerase, alanine dehydrogenase, alanopine dehydrogenase, alcohol dehydrogenase, alcohol dehydrogenase (NAD(P)+), aldehyde dehydrogenase (NAD(P)+), aldehyde (NAD+) dehydrogenase, aldose 1-dehydrogenase, alkene monooxygenase, alpha-santonin 1,2-reductase, aminobutyraldehyde dehydrogenase, aminomuconate-semialdehyde dehydrogenase, anthocyanide reductase, anthranilate 1,2-dioxygenase (deamination, decarboxylation), anthraniloyl-CoA monooxygenase, apiose 1-reductase, aquacobalamin reductase, arogenase, arogenase (NAD(P)+), aryl-alcohol dehydrogenase, aryl-aldehyde dehydrogenase, asparagusate reductase,aspartate dehydrogenase, ATP-dependent NAD(P)H-hydrate hydratase, benzaldehyde dehydrogenase (NAD+), benzene 1,2-dioxygenase, benzoate 1,2-dioxygenase, beta-alanopine dehydrogenase, betaine, Petition 870250049834, dated 06 / 13 / 2025, page 51 / 97 43 / 76 aldehyde dehydrogenase, biphenyl 2,3-dioxygenase, butanal dehydrogenase, carnitine 3-dehydrogenase, CDP-4-dehydro-6-deoxyglucose reductase, CDP-glucose 4,6-dehydratase, CDP-paratose 2-epimerase, cholestylene-5-ene-3-beta,7-alpha-diol 3-beta-dehydrogenase, cholestylenetetraol 26-dehydrogenase, cis-1,2-dihydro-1,2-dihydroxynaphthalene dehydrogenase, cis-1,2-dihydrobenzene-p-1,2-diol dehydrogenase, cis-1,2-dihydroxy-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-dihydroethylcatechol dehydrogenase, CoAbisulfite reductase, Ca ob(II)alamine reductase, coniferyl aldehyde dehydrogenase, cucurbitacin delta23-reductase, cyclohexane-1,2-diol dehydrogenase, cyclohexanol dehydrogenase, cyclopentanol dehydrogenase, cystine reductase, D-arabinitol 2-dehydrogenase, D-arabinitol 4-dehydrogenase, D-arabinose 1-dehydrogenase,Darabinose 1-dehydrogenase (NAD(P)+), D-iditol 2-dehydrogenase, D-malate dehydrogenase (decarboxylation), D-threo-aldose 1-dehydrogenase, D-xylose 1-dehydrogenase, D-xylulose reductase, dihydrodiol dibenzothiophene dehydrogenase, diferric transferrin reductase, dihydrouracil dehydrogenase (NAD+), diiodophenylpyruvate reductase, dimethyl malate dehydrogenase, DTDP-glucose 4,6-dehydratase, ephedrine dehydrogenase, erythrose-4-phosphate dehydrogenase, estradiol 17-alpha dehydrogenase, estradiol 17-beta dehydrogenase, acyl-CoA fatty synthase, ferredoxin-NAD(+) reductase, ferric chelate reductase, fluoren-9-ol dehydrogenase, fluoroacetaldehyde dehydrogenase, FMN reductase, formaldehyde dehydrogenase, fruccurnate reductase, fumarate reductase (NADH), fulfurylfuramide 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, Petition 870250049834, dated 06 / 13 / 2025, page 52 / 97 44 / 76 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, kinurenate-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-gulonate 3-dehydrogenase, L-iditol 2-dehydrogenase, lidonate 5-dehydrogenase, L-rhamnose 1-dehydrogenase, L-threonate 3-dehydrogenase, L-threonine 3-dehydrogenase, lactaldehyde dehydrogenase, lactaldehyde reductase, latosterol oxidase, leucohemoglobin reductase, leucine dehydrogenase, long-chain alcohol dehydrogenase, lysine dehydrogenase, malate dehydrogenase (decarboxylation), malate dehydrogenase (oxaloacetate decarboxylation), maleyl acetate reductase,malonate semialdehyde dehydrogenase, malonate semialdehyde dehydrogenase (acetylation), mannitol 2-dehydrogenase, mannitol dehydrogenase, mannitol-1-phosphate 5-dehydrogenase, mannuronate reductase, Petition 870250049834, dated 06 / 13 / 2025, page 53 / 97 45 / 76 melilotate 3-monooxygenase, meso-tartrate dehydrogenase, methanol dehydrogenase, methylenetetrahydrofolate dehydrogenase (NAD+), methylglyoxal reductase (NADH-dependent), methylmalonate semialdehyde dehydrogenase (acylation), mevaldate reductase, monodehydroascorbate reductase (NADH), morphine 6-dehydrogenase, formaldehyde-dependent mycothiol dehydrogenase, mycothione reductase, myristoyl-CoA 11-11-(E) desaturase, myristoyl-CoA 11-11-(Z) desaturase, N-acetylhexosamine 1-dehydrogenase, N-acylmannosamine 1-dehydrogenase, a N-hydroxy-2-acetamidofluorene reductase, NAD(+)-dinitrogen reductase, ADP-D-ribosyltransferase, NAD(+)-diphthamide ADP-ribosyltransferase, NAD(P)(+)-arginine protein 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, NAD+ synthase (glutamine hydrolysis), NADH dehydrogenase (quinone), NADH peroxidase, naphthalene 1,2-dioxygenase, nicotinamide nucleotide adenylyl transferase, nitric oxide dioxygenase, nitrite reductase (NAD(P)H), nitroquinoline-N-oxide reductase, octanol dehydrogenase, omega-hydroxydecanoate dehydrogenase, opino dehydrogenase, orcinol 2-monooxygenase, ornithine cyclodeaminase, orotate reductase (NADH), oxaloglycolate reductase (decarboxylation), pantoate 4-dehydrogenase, perylyl alcohol dehydrogenase, phenylacetaldehyde dehydrogenase, phenylalanine dehydrogenase, a phenyl glyoxylate dehydrogenase (acylation), phosphatidylcholine 12-monooxygenase, phosphatidylcholine desaturase, phosphogluconate 2-dehydrogenase, phosphoglycerate dehydrogenase, phosphonate dehydrogenase, phthalate 4,5-cis-dihydrodiol dehydrogenase, phthalate 4,5-dioxygenase, pimeloyl-CoA dehydrogenase, precorrin-2 dehydrogenase, precorrin-3B synthase,prephenate dehydrogenase, propanediol phosphate dehydrogenase, protein disulfide reductase, pyridoxal 4-dehydrogenase, pyrroline 2-carboxylate reductase, pyrroline 5- Petition 870250049834, dated 06 / 13 / 2025, page 54 / 97 46 / 76 carboxylate reductase, quinate dehydrogenase, retinal dehydrogenase, retinol dehydrogenase, ribitol 2-dehydrogenase, ribitol-5-phosphate 2-dehydrogenase, rubredoxin-NAD(+) reductase, rubredoxin-AD(P)(+) reductase, S-(hydroxymethyl)glutathione dehydrogenase, saccharin (NAD+, L-glutamate formation), saccharin (NAD+, L-lysine formation), salicylaldehyde dehydrogenase, salicylate 1-monooxygenase, sequoiitol dehydrogenase, serine 2-dehydrogenase, Sn-glycerol-1-phosphate dehydrogenase, sorbitol-6-phosphate 2-dehydrogenase, steroid 17-alpha-mono-oxygenase, sterol-alpha-carboxylate 3-dehydrogenase (decarboxylation), strombin dehydrogenase, succinate semialdehyde dehydrogenase, succinate-semialdehyde-semialdehyde dehydrogenase (NAD(P)+), succinyl-glutamate-semialdehyde dehydrogenase, sulcatone reductase, tagaturonate reductase, tartrate dehydrogenase, tauropine dehydrogenase, taxifolin 8-mono-oxygenase, terephthalate 1,2-cis-dihydrodiol dehydrogenase, terephthalate 1,2-dioxygenase, testosterone 17-beta-dehydrogenase, tetrahydroxypteridine cycloisomerase, thiomorpholine carboxylate dehydrogenase, TM0436, toluene dioxygenase, trans-2-enoyl-CoA reductase (NAD+), trimethylamine-N-oxide reductase, tryptophan dehydrogenase, UDP-Glucose 4-epimerase, UDP-Glucose 6-dehydrogenase, UDP-Glucuronate 5'-epimerase, UDP-Glucuronate decarboxylase, UDP-N-acetylglucosamine 6-dehydrogenase, ureidoglycolate dehydrogenase, uronate dehydrogenase, vanylate monooxygenase, a vanillin dehydrogenase, vomifoliol dehydrogenase, xanthine dehydrogenase, xanthomatin reductase, or xanthonin dehydrogenase.

[0131] In some embodiments of the present invention, the analyte-specific enzyme includes a flavin oxidase, such as a flavin adenine dinucleotide (FAD)-dependent or flavin mononucleotide (FMN)-dependent oxidase. Examples of (FAD)-dependent or (FMN)-dependent oxidases include: (R)-6-hydroxynicotine oxidase, (S)-2-hydroxyacid oxidase, (S)-6 Petition 870250049834, dated 06 / 13 / 2025, p. 55 / 97 47 / 76 hydroxynicotine oxidase, 2-enoate reductase, 2-methyl-branched-chain enoyl-CoA reductase, 2-nitropropane dioxygenase, 2,4-dichlorophenol 6-monooxygenase, 2,6-dihydroxypyridine 3-monooxygenase, 3-acyl-nitropropanoate oxidase, 3-hydroxy-2-methylpyridine carboxylate dioxygenase, 3-hydroxybenzoate 4-monooxygenase, 3-hydroxybenzoate 6-monooxygenase, 3-hydroxyphenyl acetate 6-hydroxylase, 4-aminobenzoate 1-monooxygenase, 4-cresol dehydrogenase (hydroxylation), 4-hydroxybenzoate 1-hydroxylase, 4-hydroxy benzoate 3-monooxygenase, 4-hydroxybenzoate 3-monooxygenase (NAD(P)H), 4-hydroxymandelate oxidase, 4-hydroxyphenyl acetate 1-monooxygenase, 4-hydroxyphenyl acetate 3-monooxygenase, 4-nitrophenol 2-monooxygenase, 4-sulfobenzoate 3,4-dioxygenase, 5-pyridoxate 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, beta-cyclopiazonate dehydrogenase, cellobiose dehydrogenase (acceptor), choline oxidase, CoA-glutathione reductase, cob(II)alanine reductase, cyanocobalamin reductase (cyanide removal), cyclohexylamine oxidase, D-2-hydroxy acid dehydrogenase, D-amino acid oxidase, D-arabinono-1,4-lactone oxidase, oxidase, D-glutamate (D-aspartate) oxidase, D-lactate dehydrogenase (cytochrome), D-sorbitol (acceptor) dehydrogenase, dehydrogluconate dehydrogenase, a deoxyribodipyrimidine 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, midazole acetate 4-monooxygenase, indole 2,3-dioxygenase, indole-3-acetaldehyde oxidase, isovaleryl-CoA dehydrogenase, Petition 870250049834, dated 06 / 13 / 2025, page 56 / 97 48 / 76 kynurenine 3-monooxygenase, L-amino acid oxidase, L-aspartate oxidase, L-galactonolactone oxidase, L-glutamate oxidase, L-lactate dehydrogenase (cytochrome), lactate 2-monooxygenase, latosterol oxidase, Latia-luciferin monooxygenase (demethylation), long-chain acyl-CoA dehydrogenase, lysine 2-monooxygenase, malate dehydrogenase (quinone), malate oxidase, mandelonitryl lyase, melylolate 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-acetylation), retinal dehydrogenase, rubredoxin-NAD(+) reductase, salicylate 1-monooxygenase, sarcosine dehydrogenase, short-chain acyl-CoA dehydrogenase, spermidine dehydrogenase, steroid 9-alpha-monooxygenase, tartronate semialdehyde synthase, taxifolin 8-monooxygenase, thiamine oxidase, trypanothione bisulfite reductase, UDP-N-acetyl muramate dehydrogenase, or vanillyl alcohol oxidase. Sensor Membrane

[0132] In some embodiments of the present invention, with reference to Figures 13 and 14B to 14D, the sensor 500 or a portion of the sensor 500 includes an outer membrane 520 or 335 that overlaps at least the operating electrode 501 or 320 and a sensing element 322 or a sensing area 509. Electrochemical sensors are frequently coated with an outer membrane. Petition 870250049834, dated 06 / 13 / 2025, p. 57 / 97 49 / 76 520 or 335 (for example, a polymer membrane) in order to impart stability to the detection reagents (for example, the analyte-specific enzyme 323 and the redox mediator 324), as well as to impose limitations on bulk transport, biocompatibility, and / or to prevent the electrode from becoming soiled.

[0133] In some embodiments of the present invention, the membrane is composed of two components, a hydrophilic polymer (water acceptor) and a crosslinker. The crosslinker links the polymer molecules to each other and anchors them to the sensor's detection layer. For analytes such as glucose, which are found in vivo at concentrations of about 5 mM, a flow-limiting membrane is necessary to prevent the electrode from becoming fouled. Examples of sensor flow-limiting membranes are disclosed, for example, in Mao et al. U.S. Patent No. 6,932,894, the entire content of which is incorporated herein by reference.

[0134] For analytes at lower concentrations, a flow-limiting membrane can be used with a longer accumulation time, for example, up to 30 minutes. Alternatively, for analytes at lower concentrations, a high-permeability membrane can be used in order to maintain the natural flow of the analyte to the detection layer, while also having a membrane to increase the biocompatibility of the sensor. For example, a hydrophilic membrane surface does not aggravate the body's immune system, thus reducing the risk of inflammation and other responses that could compromise sensor performance. Analyte Monitoring Systems

[0135] Therefore, the embodiments include analyte monitoring devices and systems that include an analyte sensor, at least a portion of which can be positioned below the user's skin surface for in vivo detection of an analyte in a body fluid. The systems of Petition 870250049834, dated 06 / 13 / 2025, p. 58 / 97 50 / 76 Analyte monitoring is disclosed in Word et al. (US Patent No. 6,134,461) and Hoss et al. (US Patent Application Publication No. 2012 / 0150005), the full contents of which are incorporated herein by reference. Embodiments of the present invention include fully implantable analyte sensors and analyte sensors in which only a portion of the sensor is positioned under the skin and a portion of the sensor resides above the skin, for example, for contact with a sensor control unit (which may include a transmitter), a receiving / displaying unit, a transceiver, a processor, etc. The sensor may, for example, be positioned subcutaneously in a user for continuous or periodic monitoring of an analyte level in the user's interstitial fluid. For the purposes of this description, continuous monitoring and periodic monitoring will be used interchangeably unless otherwise indicated.The sensor response can be correlated and / or converted to analyte levels in blood or other fluids. In certain embodiments, an analyte sensor can be positioned in contact with interstitial fluid to detect the analyte level, which can be used to infer the analyte level in the user's bloodstream. Analyte sensors can be inserted into a vein, an artery, or another fluid-containing body part. In some embodiments, analyte sensors can be configured to monitor the analyte level for a period of time that can range from seconds, minutes, hours, days, weeks, to months, or longer.

[0136] In some embodiments of the present invention, the analyte sensors have the ability to detect an analyte in vivo for one hour or more, for example, several hours or more, for example, several days or more, for example, three days or more, for example, five days or more, for example, seven days or more, for example, several weeks or more, or one month or more. Future analyte levels can be predicted based on the information obtained, for example, the current level of Petition 870250049834, dated 06 / 13 / 2025, p. 59 / 97 51 / 76 analyte at time t, the rate of change of analyte, etc. Predictive alarms can notify the user of a predicted analyte level that may be of interest before the user's analyte level reaches the predicted future analyte level. This provides the user with an opportunity to take corrective action.

[0137] Figure 15 shows a data monitoring and management system such as, for example, an analyte monitoring system 400 according to certain embodiments of the present invention. Aspects of the embodiments of the present invention are described primarily with respect to glucose monitoring devices and systems, and glucose detection methods, for convenience only, and such description does not in any way limit the scope of the embodiments. It should be understood that the analyte monitoring system can be configured to monitor a variety of analytes as disclosed herein simultaneously or at different times.

[0138] Analytes that can 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 hormones, hormones, 3-hydroxybutyrate, ketones, ketone bodies, peroxide, prostate-specific antigen, prothrombin, RNA, thyroid-stimulating hormone, and troponin. Analytes also include drugs such as, for example, antibiotics (e.g., gentamicin, vancomycin, and others), and digitoxin, digoxin, drugs of abuse, theophylline, and warfarin can also be monitored. In some embodiments, more than one analyte is monitored, and the analytes can be monitored simultaneously or at different times. Petition 870250049834, dated 06 / 13 / 2025, p. 60 / 97 52 / 76

[0139] The analyte monitoring system 400 includes an analyte sensor 401, a data processing unit 402 connectable to the sensor 401, and a primary receiving unit 404. In some examples, the primary receiving unit 404 is configured to communicate with the data processing unit 402 via a communication link 403. In certain embodiments, the primary receiving unit 404 may also be configured to transmit data to a data processing terminal 405 to evaluate or otherwise process or format the data received by the primary receiving unit 404. The data processing terminal 405 may be configured to receive data directly from the data processing unit 402 via a communication link 407, which may optionally be configured for bidirectional communication.In addition, the data processing unit 402 may include a transmitter or a transceiver for transmitting and / or receiving data to and / or from the primary receiving unit 404 and / or the data processing terminal 405 and / or optionally a secondary receiving unit 406.

[0140] Figure 15 also shows an optional secondary receiving unit 406 that is operatively coupled to the communication link 403 and configured to receive data transmitted from the data processing unit 402. The secondary receiving unit 406 can be configured to communicate with the primary receiving unit 404 as well as the data processing terminal 405. In some embodiments, the secondary receiving unit 406 can be configured for bidirectional wireless communication with either the primary receiving unit 404 or the data processing terminal 405. As discussed in detail below, in some instances, the secondary receiving unit 406 may be a less-characterized receiver compared to the primary receiving unit 404; for example, the secondary receiving unit 406 may include a limited or minimal number of functions and Petition 870250049834, dated 06 / 13 / 2025, pp. 61 / 97 53 / 76 features compared to the primary receiving unit 404. Thus, the secondary receiving unit 406 may include a smaller compact enclosure (in one or more dimensions, including all), or be incorporated into a device that includes a wristwatch, an armband, a PDA, an MP3 player, a mobile phone, etc., for example. Alternatively, the secondary receiving unit 406 may be configured with identical and substantially similar functions and features to those of the primary receiving unit 404. The secondary receiving unit 406 may include a docking portion configured to couple with a mounting splice unit for placement, for example, on a nightstand for nighttime monitoring, and / or a two-way communication device. A mounting splice may carry a power supply.

[0141] Only one analyte sensor 401, one data processing unit 402, and one data processing terminal 405 are shown in the embodiment of the analyte monitoring system 400 illustrated in Figure 15. However, it should be appreciated by one normally versed in the state of the art that the analyte monitoring system 400 may include more than one sensor 401 and / or more than one data processing unit 402, and / or more than one data processing terminal 405. Multiple sensors may be positioned on a user for analyte monitoring at the same time or at different times. In certain embodiments, the analyte information obtained by a first sensor positioned on a user may be used as a comparison with the analyte information obtained by a second sensor. This may be useful to confirm or validate the analyte information obtained from one or both sensors.Such redundancy can be useful if analyte information is considered in decisions related to critical therapies. In certain modalities, a first sensor can be used to calibrate a second sensor.

[0142] The analyte 400 monitoring system can be a system of Petition 870250049834, dated 06 / 13 / 2025, page 62 / 97 54 / 76 continuous, or semi-continuous, monitoring, or a distinct monitoring system. In a multi-component environment, each component can be configured to be uniquely identified by one or more of the other components in the system so that a communication conflict can be resolved immediately between the various components within the analyte 400 monitoring system. For example, unique IDs, communication channels, and others can be used.

[0143] In certain embodiments, the sensor 401 is physically positioned on or over the body of a user whose analyte level is being monitored. The sensor 401 may be configured to at least periodically sample the user's analyte level and convert the sampled analyte level into a corresponding signal for transmission by the data processing unit 402. The data processing unit 402 may be coupled to the sensor 401 so that both devices are positioned on or over the user's body, with at least a portion of the analyte sensor 401 positioned transcutaneously. The data processing unit may include a fastening element, such as an adhesive or other means, to secure it to the user's body. An accessory attachable to the user and attachable to the data processing unit 402 may be used. For example, an accessory may include an adhesive surface.The data processing unit 402 performs data processing functions, which may include, but are not limited to, filtering and encoding of data signals, each of which corresponds to a sampled level of the user's analyte, for transmission to the primary receiving unit 404 via the communication link 403. In some embodiments, the sensor 401 or the data processing unit 402 or a combined sensor / data processing unit may be fully implantable under the surface of the user's skin.

[0144] In certain modes, the primary receiving unit 404 Petition 870250049834, dated 06 / 13 / 2025, pp. 63 / 97 55 / 76 may include an analog interface section including an RF receiver and an antenna that is configured to communicate with the data processing unit 402 via the communication link 403, and a data processing section to process the data received from the data processing unit 402 including data decoding, error detection and correction, data trigger pulse generation, data bit recovery, etc., or any combination thereof.

[0145] In operation, the primary receiving unit 404 in certain embodiments is configured to synchronize with the data processing unit 402 to uniquely identify the data processing unit 402, based, for example, on an identification information of the data processing unit 402, and then periodically receive the signals transmitted from the data processing unit 402 associated with the monitored analyte levels detected by the sensor 401.

[0146] Referring again to Figure 15, the data processing terminal 405 may include a personal computer, a portable computer including a laptop or handheld device (e.g., a personal digital assistant (PDA)), a telephone including a mobile phone (e.g., a multimedia- or internet-enabled mobile phone including an iPhone™, a Blackberry®, or a similar phone), an mp3 player (e.g., an iPod™, etc.), a pager, and others), and / or a drug delivery device (e.g., an infusion device), each of which may be configured for data transmission with the receiver via a wired or wireless connection. Furthermore, the data processing terminal 405 may also be connected to a data network (not shown) to store, retrieve, update, and / or analyze data corresponding to the user's detected analyte level.

[0147] The 405 data processing terminal may include a Petition 870250049834, dated 06 / 13 / 2025, p. 64 / 97 56 / 76 drug delivery device (e.g., an infusion device), such as an insulin infusion pump or similar, which can be configured to administer a drug (e.g., insulin) to the user, and which can be configured to communicate with the primary receiving unit 404 for receiving, among other things, the measured analyte level. Alternatively, the primary receiving unit 404 can be configured to integrate an infusion device so that the primary receiving unit 404 is configured to administer an appropriate drug (e.g., insulin) to users, for example, to administer and modify baseline profiles, as well as to determine appropriate boluses for administration based, among other things, on the detected analyte levels received from the data processing unit 402.An infusion device can be an external device or an internal device, such as a device that is fully implantable in a user.

[0148] In certain embodiments, the data processing terminal 405, which may include an infusion device, for example, an insulin pump, may be configured to receive analyte signals from the data processing unit 402, and thus incorporates the functions of the primary receiving unit 404, including data processing to control the user's insulin therapy and analyte monitoring. In certain embodiments, the communication link 403, as well as one or more of the other communication interfaces shown in Figure 15, may use one or more wireless communication protocols such as, but not limited to: an RF communication protocol, an infrared communication protocol, a Bluetooth-enabled communication protocol, an 802 wireless communication protocol.11x, or an equivalent wireless communication protocol that allows secure wireless communication between multiple units (e.g., according to the requirements of the Health Insurance Portability and Accountability Act (HIPPA)), while avoiding collision and... Petition 870250049834, dated 06 / 13 / 2025, p. 65 / 97 57 / 76 potential data interference.

[0149] In other embodiments, the data processing unit 402 and / or the primary receiving unit 404 and / or the secondary receiving unit 406, and / or the data processing terminal (infusion device) 405 may be configured to wirelessly receive the analyte value via a communication link, for example, a blood analyte meter. In other embodiments, a user manipulating or using the analyte monitoring system 400 (FIGURE 15) may manually enter the analyte value by using, for example, a user interface (e.g., a keyboard, a keypad, voice commands, and others) incorporated into one or more of the data processing unit 402, the primary receiving unit 404, the secondary receiving unit 406, or the data processing terminal (infusion device) 405.

[0150] A sensor (e.g., an enzymatic biosensor) as disclosed herein for measuring low nanomolar concentrations of an analyte may be used in an in vivo monitoring system and, while positioned in vivo on a user (e.g., a human individual), comes into contact with the user's bodily fluid and detects one or more levels of analyte contained therein. An in vivo monitoring system may include one or more reader devices that receive the detected analyte data from a sensor control device. These reader devices may process and / or display the detected analyte data, or sensor data, in any number of forms, to the user.

[0151] With reference to Figure 16, in some embodiments, a reading device 120 may be a mobile communication device, such as a dedicated reading device (configured for communication with a sensor control device 102 (FIGURE 17), and optionally a computerized system, but without telephony communication capability) or a Petition 870250049834, dated 06 / 13 / 2025, pp. 66 / 97 58 / 76 mobile phone including, but not limited to, a Wi-Fi or Internet-enabled smartphone, a tablet, or a personal digital assistant (PDA). Examples of smartphones may include mobile phones based on a Windows® operating system, an Android™ operating system, an iPhone® operating system, a Palm® WebOS™ operating system, a Blackberry® operating system, or a Symbian® operating system, with data network connectivity functionality for data transmission over an Internet connection and / or a local area network (LAN).

[0152] A reading device 120 can also be configured as a mobile smart wearable electronic set, such as an optical set that is worn over or adjacent to the user's eye (e.g., smart glasses or smart goggles, such as Google Glasses, which is a mobile communication device). This optical set may have a transparent display that shows information about the user's analyte level to the user while at the same time allowing the user to see through the display in such a way that the user's overall vision is minimally obstructed. The optical set may have wireless communication capabilities similar to a smartphone. Other examples of wearable electronics include devices that are worn around or in close proximity to the user's wrist (e.g., a watch, etc.), neck (e.g., a necklace, etc.), head (e.g., a headband, a hat, etc.), chest, or elsewhere.

[0153] Figure 16 is a block diagram of an exemplary embodiment of a reading device 120 configured as a smartphone. Here, the reading device 120 includes an input component 121, the dial 122, and the processing circuits 206, which may include one or more processors, microprocessors, controllers and / or microcontrollers, each of which may be a distinct chip or distributed among (and a portion of) a number Petition 870250049834, dated 06 / 13 / 2025, pp. 67 / 97 59 / 76 of different chips. Here, the processing circuit 206 includes a communications processor 202 which has on-board memory 203 and an application processor 204 which has on-board memory 205. The reading device 120 also includes the RF communication circuit 208 coupled with an RF antenna 209, a memory 210, a multifunctional circuit 212 with one or more associated antennas 214, a power supply 216, a power management circuit 218, and a clock 219. Figure 16 is a summary representation of the typical hardware and functionality that resides within a smartphone, and those normally versed in the state of the art will immediately recognize that other hardware and functionalities (e.g., codecs, drivers, pasting logic) may also be included.

[0154] Figure 16 also shows that the communications processor 202 can interface with the RF communication circuit 208 and perform analog-to-digital conversions, encoding and decoding, digital signal processing, and other functions that facilitate the conversion of voice, video, and data signals into a format (e.g., phase and quadrature) suitable for provision to the RF communication circuit 208, which can then wirelessly transmit the signals. The communications processor 202 can also interface with the RF communication circuit 208 to perform the reverse functions necessary to receive a wireless transmission and convert it into digital data, voice, and video. The RF communication circuit 208 may include a transmitter and a receiver (e.g., integrated as a transceiver) and the associated encoder logic.

[0155] Referring again to Figure 16, the application processor 204 can be adapted to run the operating system and any software applications residing on the reading device 120, process video and graphics, and perform those other functions unrelated to processing communications transmitted and received through the RF antenna 209. The system Petition 870250049834, dated 06 / 13 / 2025, pp. 68 / 97 The smartphone's operating system will operate in conjunction with a number of applications on the reading device. Any number of applications (also known as user interface applications) can run on the reading device at any given time, and may include one or more applications that are related to a diabetes monitoring regimen, in addition to other commonly used applications that are not related to such a regimen, for example, email, calendar, weather forecast, sports, games, etc. For example, data indicating a detected analyte level and in vitro blood analyte measurements received by the reading device can be securely communicated to the user interface applications residing in the memory of the reading device. Such communications can be securely performed, for example, through the use of mobile application containerization or overlay technologies.

[0156] Memory 210 may be shared by one or more of the various functional units present within the reading device 120, or it may be distributed between two or more of them (for example, as separate memories present within different chips). Memory 210 may also be a separate chip from itself. Memories 203, 205, and 210 are non-transient, and may be volatile memories (e.g., RAM, etc.) and / or non-volatile memories (e.g., ROM, flash memory, F-RAM, etc.).Multifunctional circuits 212 can be implemented as one or more chips and / or components (e.g., transmitter, receiver, transceiver, and / or other communication circuit) that perform other functions such as local wireless communications, for example, with the sensor control device 102 under the appropriate protocol (e.g., Wi-Fi, Bluetooth, Bluetooth Low Energy, Near Field Communication (NFC), Radio Frequency Identification (RFID), proprietary protocols, and others) and determine the geographic location of the reading device 120 (e.g., a global positioning system (GPS) hardware). One or more other antennas 214 are. Petition 870250049834, dated 06 / 13 / 2025, pp. 69 / 97 61 / 76 associated with the functional circuits 212 as needed to operate with the various protocols and circuits.

[0157] The power supply 216 may include one or more batteries, which may be rechargeable or single-use disposable batteries. The power management circuit 218 may regulate battery charging and power supply monitoring, boost power, perform DC conversions, and more.

[0158] The 120 reading device may also include or be integrated with a drug delivery device (e.g., insulin, etc.) in such a way that, for example, they share a common housing. Examples of such drug delivery devices may include medication pumps that have a cannula that remains in the body to allow infusion over a period of multiple hours or multiple days (e.g., wearable pumps for basal and bolus insulin delivery). The 120 reading device, when combined with a medication pump, may include a reservoir to store the drug, a pump connectable to the transfer tubing, and an infusion cannula. The pump may force the drug out of the reservoir, through the tubing and into the diabetic's body via the cannula inserted therein.Other examples of drug delivery devices that can be included (or integrated with) the 120 reading device include portable injection devices that pierce the skin only for each application and are subsequently removed (e.g., insulin pens). A 120 reading device, when combined with a portable injection device, may include an injection needle, a cartridge to load the drug, an interface to control the amount of drug to be administered, and an actuator to cause the injection to occur. The device can be used repeatedly until the drug is depleted, at which point the combined device can be discarded, or the cartridge can be. Petition 870250049834, dated 06 / 13 / 2025, pp. 70 / 97 62 / 76 can be replaced with a new one, at which point the combined device can be reused repeatedly. The needle can be replaced after each injection.

[0159] The combined device can 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 circuit system that rarely requires user intervention to operate, such as to confirm dose changes). For example, the diabetic's analyte level can be monitored automatically and repeatedly by the sensor control device 102, which can then communicate this monitored analyte level to the reading device 120, and the appropriate drug dosage to control the diabetic's analyte level can be automatically determined and subsequently applied to the diabetic's body.Software instructions for controlling the pump and the amount of insulin administered can be stored in the memory of the reading device 120 and executed by the processing circuit of the reading device. These instructions can also cause the calculation of drug administration amounts and durations (e.g., a bolus infusion profile and / or a basal infusion profile) based on analyte level measurements obtained directly or indirectly from the sensor control device 102. In some embodiments, the sensor control device 102 can determine the drug dosage and communicate it to the reading device 120.

[0160] Figure 17 is a block diagram illustrating an exemplary embodiment of the sensor control device 102 having the analyte sensor 104 and the sensor electronic components 250 (including the analyte monitoring circuits) which may have most of the processing capacity to make the final result data suitable for display to the user. In Figure 17, a single semiconductor chip 251 is shown, which may be a custom application-specific integrated circuit (ASIC). [The remaining text appears to be incomplete and possibly contains errors.] Petition 870250049834, dated 06 / 13 / 2025, pp. 71 / 97 63 / 76 within the ASIC 251 certain high-level functional units, including an analog front end (AFE) 252, a power management (or control) circuit 254, a processor 256, and a communication circuit 258 (which may be implemented as a transmitter, a receiver, a transceiver, a passive circuit, or otherwise according to the communication protocol). In this embodiment, the AFE 252 and the processor 256 are used as analyte monitoring circuits, but in other embodiments either circuit 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 a distinct chip or distributed among (and a portion of) a number of different chips.

[0161] A memory 253 can also be included within the ASIC 251 and can be shared by the various functional units present within the ASIC 251, or it can be distributed between two or more of them. The memory 253 can also be a separate chip. The memory 253 is non-transient and can be volatile and / or non-volatile memory. In this embodiment, the ASIC 251 is coupled with the power supply 260, which can be a coin cell battery, or another type. The AFE 252 interfaces with the in vivo analyte sensor 104 and receives the measurement data from it and sends the data to the processor 256 in digital form, which, in turn, in some embodiments, can process it in any appropriate way. This data can then be provided to the communication circuit 258 to send, via the antenna 261, to the reading device 120, for example, where minimal additional processing is required by the resident software application to display the data.The 261 antenna can be configured according to the application requirements and communication protocol. The 261 antenna can be, for example, a PCB trace antenna, a ceramic antenna, or a distinct metallic antenna. The 261 antenna can be configured as an antenna. Petition 870250049834, dated 06 / 13 / 2025, pp. 72 / 97 64 / 76 single-pole antenna, a two-pole antenna, an F-type antenna, a loop antenna, and others.

[0162] Information can be communicated from sensor control device 102 to a second device (e.g., reading device 120) on the initiative of sensor control device 102 or reading device 120. For example, information can be communicated automatically and / or repeatedly (e.g., continuously) by sensor control device 102 when analyte information is available, or according to a schedule (e.g., every 1 minute, every 5 minutes, every 10 minutes, or others), in which case the information can be stored or logged in a memory of sensor control device 102 for later communication. Information can be transmitted from sensor control device 102 in response to receiving a request from the second device.This request can be an automated request, for example, a request transmitted by the second device according to a schedule, or it can be a request generated by a user's initiative (for example, an ad hoc or manual request). In some embodiments, a manual request for data is indicated as a scan of the sensor control device 102 or an on-demand data transfer from device 102. In some embodiments, the second device can transmit a packet of voting signals or data to the sensor control device 102, and device 102 can treat each vote (or votes that occur at certain time intervals) as a request for data and, if the data is available, then transmit such data to the second device.In many embodiments, communication between sensor control device 102 and the second device is secure (e.g., encrypted and / or between authenticated devices), but in some embodiments data can be transmitted from sensor control device 102 in an insecure manner, for example, as a transmission to all devices. Petition 870250049834, dated 06 / 13 / 2025, pp. 73 / 97 65 / 76 listening on the track.

[0163] Different types and / or forms and / or amounts of information may be sent as part of each communication including, but not limited to, one or more of the current sensor measurements (e.g., the most recently obtained information on the analyte level that corresponds temporally to the moment at which the reading is initiated), the rate of change of the measured metric over a predetermined period of time, the rate of change of the metric information (acceleration in the rate of change), or historical metric information that corresponds to metric information obtained before a given reading and stored in a memory of the sensor control device 102.

[0164] Some or all of the information about real-time, history, rate of change, rate of change ratio (such as acceleration or deceleration) can be sent to the reading device 120 in a given communication or transmission. In certain embodiments, the type and / or form and / or amount of information sent to the reading device 120 can be pre-programmed and / or unchangeable (e.g., pre-configured at manufacturing), or it can be not pre-programmed and / or unchangeable so that it can be selected and / or changed in the field one or more times (e.g., by activating a system switch, etc.).Therefore, in certain embodiments the reading device 120 can send (in real time) an analyte value derived from the sensor (e.g., in numerical format), a current rate of change of the analyte (e.g., in the form of an analyte rate indicator such as an arrow pointing in a direction to indicate the current rate), and historical analyte trend data based on sensor readings acquired by and stored in the memory of the sensor control device 102 (e.g., in the form of a graphic trace). Furthermore, a reading or measurement of skin or sensor temperature can be collected by an optional temperature sensor. Petition 870250049834, dated 06 / 13 / 2025, pp. 74 / 97 66 / 76 257. These readings or measurements can be communicated (either individually or as an aggregate measurement over time) from the sensor control device 102 to another device (e.g., a reader or the reading device 120). The temperature reading or measurement, however, can be used in conjunction with a software routine executed by the reading device 120 to correct or compensate for the analyte measurement sent to the user, instead of or in addition to actually displaying the temperature measurement to the user.

[0165] The following examples are presented for illustrative purposes only and do not limit the scope or content of this patent application. EXAMPLES Example 1. Calculation of the accumulation sensitivity mode detection when using polymer-coated sensors and long accumulation times.

[0166] Figure 5 shows the calibration curves obtained through amperometry and accumulation mode detection when 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 amperometry, accumulation mode detection allows the sensor sensitivity to be easily adjusted by changing the accumulation time. For peak height and peak area measurements, the sensor sensitivity is increased by a factor of approximately 10 by increasing the accumulation time from 1 minute to 10 minutes. The sensitivity for each calibration curve shown in Figure 5 was calculated as the slope of the linear fit with the tabulated data shown in Table 1. Table 1 Accumulation time (minutes) Sensitivity Amperometry (nA / μM) Peak height of accumulation mode (nA / μM) Peak area of ​​accumulation mode (nC / μM) 1 0.0022 0.0043 0.11 Petition 870250049834, dated 06 / 13 / 2025, pp. 75 / 97 67 / 76 2 0.0023 0.0086 0.26 5 0.0024 0.020 0.65 10 0.0025 0.039 1.33

[0167] Since peak height and amperometry measurements are made in the same units, their sensitivities can be directly compared. Using the flow membrane sensor data as shown in Figure 5, the ratio (i.e., time increase) between the accumulation mode sensitivity and the amperometry sensitivity under equivalent sensor conditions was calculated using the tabulations shown in Table 2. As indicated, at an accumulation time of 1 minute, the sensor sensitivity is 2 times greater than when using accumulation mode detection compared to amperometry. Therefore, with an increase in accumulation time to 10 minutes, the difference in sensitivity increases to 15 times. Table 2 Accumulation time (minutes) Peak height / Amperometry ratio 1 2.0 2 3.7 5 8.3 10 15.6 Example 2. Optimization of the accumulation mode signal for high-sensitivity detection with increased frequency and the addition of carbon nanotubes.

[0168] Figure 7 shows the detection of the 200 nM accumulation mode. Petition 870250049834, dated 06 / 13 / 2025, pp. 76 / 97 68 / 76 of glucose under two different signal filtering frequencies of 0.032 Hz and 3.2 Hz. As shown, the detection peak is much sharper when using the higher frequency filter, leading to a greater peak height. The area under the two curves, however, does not change. This shows that, when using peak height measurement, a higher frequency filter is ideal for maximizing signal magnitude. In particular, it was found that changing the filtering frequency from 0.032 Hz to 3.2 Hz increases the peak height signal by a factor of 2 to 3. Furthermore, with filtering frequencies higher than 3.2 Hz, the signal noise was too high to make accurate measurements of the amperometric current and peak accumulation characteristics (peak height and area).

[0169] As a means of enhancing the signal of the accumulation mode, carbon nanotubes (CNTs) were added to make the deposited detection reagent more uniform and electrically conductive, thereby increasing the kinetics of the redox-mediated oxidation step. This increase in kinetics resulted in the accumulation mode current peak having a higher peak height. Figure 8A shows micrographs of the glucose detection reagent deposited and cured with and without CNTs. As shown, the detection reagent containing CNTs is deposited more uniformly, whereas the detection reagent without CNTs exhibits a large coffee ring effect. The addition of CNTs to the detection reagent was found to increase the signal peak height by a factor of 5 to 6.

[0170] Furthermore, Figure 8B shows the results of a probing experiment of the effect of signal filtering frequency and the addition of CNTs to the detection reagent on sensor sensitivity when using amperometry and accumulation detection mode as measured by peak height and peak area when using exemplary glucose sensors at glucose concentrations from 0 to 200 nM as indicated. Four sensors of both types (with and without the CNTs in Petition 870250049834, dated 06 / 13 / 2025, pp. 77 / 97 69 / 76 detection reagents) were tested, and each calibration curve is the average response of the four indicated sensors. An accumulation time of ten minutes was used for each accumulation mode detection. Two consecutive measurements were made at each glucose concentration: one using a filtering frequency of 0.032 Hz and one using a filtering frequency of 3.2 Hz.

[0171] The sensitivity for each calibration curve in Figure 8B was calculated as the slope of the linear fit, and the tabulated data are shown in Table 3. As seen, the sensitivity of the amperometric measurement sensor changes minimally with the filtering frequency and the presence of CNTs, remaining below 0.0003 nA / nM for all conditions. For the accumulation mode measurement when using peak area, the sensor sensitivity does not change with the filtering frequency but increases slightly with the addition of CNTs to the detection reagent. The most drastic changes in sensor sensitivity are observed for the accumulation mode measurement when using peak height. The filtering frequency and the addition of CNTs to the detection reagent increase the sensor sensitivity.Increasing the filtering frequency from 0.032 Hz to 3.2 Hz increases the sensitivity by a factor of about 2.5, while adding CNTs to the detection reagent increases the sensitivity by a factor of about 5.5. Furthermore, increasing the filtering frequency combined with the addition of CNTs increases the sensitivity of the accumulation mode measurement by a factor of about 14. Table 3 Variables Sensitivity Filtering frequency (Hz) CNTs in the detection reagent? Amperometry (nA / nM) Peak height of accumulation mode (nA / nM) Peak area of ​​accumulation mode (nC / nM) Petition 870250049834, dated 06 / 13 / 2025, pp. 78 / 97 70 / 76 0.032 No 0.00023 0.0071 0.11 3.2 No 0.00024 0.018 0.10 0.032 Yes 0.00026 0.041 0.14 3.2 Yes 0.00027 0.10 0.15

[0172] Since peak height and amperometry measurements are made in the same units, their sensitivities can be directly compared. Table 4 provides the ratio between the accumulation mode sensitivity and the amperometry sensitivity under equivalent sensor conditions. As shown, even at a filtering frequency of 0.032 Hz and without CNTs in the detection reagent, the sensor sensitivity is 30 times greater when using accumulation mode detection compared to amperometry. Therefore, with increased filtering frequency and the addition of CNTs to the detection reagent to optimize the peak height of the accumulation mode, the difference in sensitivity increases to almost 400 times. Table 4 Variables Ratio Filtering Frequency (Hz) CNTs in detection reagent? Peak height / Amperometry 0.032 No 31 3.2 No 75 0.032 Yes 158 3.2 Yes 370 Petition 870250049834, dated 06 / 13 / 2025, p. 79 / 97 71 / 76 Example 3. Comparison of sensitivity, detection limit, and linear range for amperometry and accumulation detection mode when using an accumulation time of 30 minutes, a frequency of 3.2 Hz, and the addition of carbon nanotubes.

[0173] As shown in Figure 9B, the currents associated with amperometric measurements are excessively small (< 50 pA) and lose linearity below 100 nM, whereas the signals for accumulation mode detection are much larger and retain linearity well below 100 nM. Table 5 below shows the sensitivity, the lower limit of detection (LOD) (calculated as 3σ / slope, using standard approach 1), and the linear detection range associated with these measurements as disclosed in Example 5. Standard approach 1 is disclosed in Mocak et al., Pure Appl. Chem. 1997, 69:297-328, the full content of which is incorporated herein by reference. In the particular approach, standard 1 is a method for calculating LOD as 3σ / slope, where σ is the standard deviation of the mold and slope is the slope of the calibration curve. Table 5 Measurement Method Sensitivity LOD / nM Linear Range / μM Amperometry 0.00017 ± 0.00001 nA / nM 120 ± 42 0.12 - >100 Accumulation Mode Peak Height 0.14 ± 0.03 nA / nM 20 ± 16 0.02 - 2 Accumulation Mode Peak Area 0.33 ± 0.04 nC / nM 4.7 ± 1.4 0.004 - 5 Example 4. Analysis of the Background Signal.

[0174] With reference to Figures 9A and 9B, a background (cathodic) signal is observed when detection is performed on the buffer solution that is open to the atmosphere. Without being limited by any theory, the oxygen reduction reaction is likely responsible for this negative background. Specifically, the mediator of Petition 870250049834, dated 06 / 13 / 2025, pp. 80 / 97 72 / 76 osmium redox and CNTs can catalyze the oxygen reduction reaction, which should result in the oxidation of the osmium mediator and resulting in an accumulation of Os3+ when the circuit is disconnected during the accumulation period. When the circuit is reconnected, this accumulation of Os3+ can be reduced, resulting in a cathodic peak. To test this hypothesis, exemplary glucose sensors were tested in 100 mM phosphate buffer containing no glucose under atmospheric and oxygen purge conditions (e.g., by bubbling). Figure 10A shows the resulting accumulation mode signal obtained for a representative sensor for accumulation times of 2, 5, and 10 minutes under atmospheric and oxygen purge conditions, as indicated. As observed, the signals are cathodic peaks under atmospheric conditions, whereas under oxygen purge conditions the signals are smaller anodic peaks.The average signals (mean) for 4 sensors are plotted in Figure 10B. As shown, the amperometry signal is observed to be slightly negative under atmospheric conditions and slightly positive under oxygen purge conditions. The results of this experiment indicate that the negative background is due to oxygen reduction catalyzed with Os. Example 5. Linear Detection Range.

[0175] To determine the linear detection range of the accumulation mode, the calibration experiment shown in Figures 9A and 9B was performed up to glucose concentrations of 200 μM. The resulting calibration curves for amperometry and accumulation mode are shown in Figure 11. The best linear fit determined for concentrations from 0 to 200 nM was predicted for higher concentrations. As seen, the amperometry signal remains linear up to at least 100 μM. The accumulation mode signal, on the other hand, remains linear up to 2 to 5 μM before starting to plateau at higher concentrations. This should be expected, since the Os redox mediator has a Petition 870250049834, dated 06 / 13 / 2025, pp. 81 / 97 73 / 76 finite charge storage capacity. For the sensors used in this experiment, this capacity appears to be around 5,000 nC. It can be observed that the linear detection range of the accumulation mode can be shifted to higher concentrations if a shorter accumulation time is used. For the data shown in this paper, a relatively long accumulation time, for example, 30 minutes, was used to obtain high sensitivity. Example 6. Materials.

[0176] Carbon sensors printed on PET substrates were obtained from Steven Label, Inc. (Santa Fe Springs, CA). The active area of ​​the operating electrode was defined by the deposited area of ​​a glucose oxidation catalyst, which was approximately 0.1 mm2. A proprietary redox polymer used for glucose oxidase (GOx) spinning and a proprietary flow-limiting membrane polymer were synthesized according to published procedures and were obtained from Nanosin, Inc. (Santa Rosa, CA) and Regis Technologies, Inc. (Morton Grove, IL), respectively. Glucose oxidase (GOx, EC 1.1.3.4, activity of 130 U / mg) from Aspergillus sp. II was obtained from Toyobo Co., Ltd. (Osaka, Japan). Poly(ethylene glycol) diglycidyl ether (400) (PEGDGE 400) and glyceryl triglycidyl ether were obtained from Polysciences, Inc. (Warrington, PA).Multilayer carbon nanotubes (CNTs, outer diameter 20–40 nm, length 10–20 μm) were sourced from MK Nano (Mississauga, Ontario, Canada). Glucose and common chemicals used for the buffer solutions were sourced from Sigma-Aldrich (St. Louis, MO). All aqueous solutions were obtained using >18.0 MΩ.cm⁻¹ deionized water from a Thermo Scientific Barnstead E-Pure water purification system. Example 7. Sensor Manufacturing.

[0177] Two different types of glucose detection reagents were used, one without CNTs and one with CNTs. The reagent without CNTs was prepared as follows Petition 870250049834, dated 06 / 13 / 2025, pp. 82 / 97 74 / 76 follows. First, three solutions were prepared in 10 mM 4-(2-hydroxyethyl)piperazine-1-ethanesulfonic acid (HEPES) buffer (pH 8): 4% (w / v) redox polymer, 8.08% (w / v) GOx, and 8.08% (w / v) PEGDGE400. These three solutions were mixed at a ratio of 3.04:5.1:1.86 to obtain the glucose detection reagent. To prepare the glucose detection reagent with CNTs, the above procedure was followed, except that the 4% redox polymer solution and the 8.08% PEGDGE400 solution were prepared in a 5% (w / v) aqueous CNT solution instead of the 10 mM HEPES solution. After preparation, the glucose detection reagent was dispensed onto the carbon working electrode of the sensor using a microsyringe (Hamilton Co.) in 15 nl aliquots. The active area of ​​each working electrode was defined by the area of ​​the dispensed detection reagent droplet. This area was typically 0.1 mm2.After dispensing the detection reagent, sensors were cured at 25°C and 60% relative humidity for at least 12 hours. For the sensors used in the experiment shown in Figure 5, an external flow-limiting polymer membrane was applied to the sensors. This membrane, consisting of a 4:1 mixture by volume of 14% (w / volume) membrane polymer and 3.5% (w / volume) glyceryl triglycidyl ether in 80 / 20 ethanol / water, was applied by immersion coating as previously described in Liu et al., Anal. Chem. 2012, 84:3403-3409, the full text of which is incorporated herein by reference. Example 8. Electrochemical Measurements.

[0178] Unless otherwise indicated, all electrochemical measurements were made using an appropriate three-electrode cell with the glucose sensor as the working electrode, an Ag / AgCl reference electrode (in 3 M kCl; Bioanalytical Systems, Inc.), and a carbon counter electrode. Petition 870250049834, dated 06 / 13 / 2025, pp. 83 / 97 75 / 76 printed on screen. The current versus (against) time curve for a sensor was measured throughout the course of an accumulation mode experiment using a potentiostat. For an accumulation mode measurement, the operating electrode was electrically disconnected from the potentiostat for a stipulated amount of time (the accumulation time), after which it was reconnected to the circuit. Figure 2 shows a schematic of the electrode diagram. When the operating electrode of a sensor was electrically connected, it was balanced at +40 mV. For the experiments shown in Figures 3A to 3D, 4A and 4B, 5, and 6A to 6H, a BASi Petit Ampere potentiostat (model LC-3D; Bioanalytical Systems, Inc., West Lafayette, IN) was used for the current measurements. A sampling interval of 0.5 seconds (s) and a filter of 0.03 Hz were used, and the current signal was recorded using the in-house LabView software (National Instruments).For all other experiments, an increased time resolution was desired. Therefore, a potentiostat with a higher time resolution was used (model 1030C; CH Instruments, Inc., Austin, TX). This potentiostat was used with a sampling interval of 0.1 seconds and a 3.2 Hz filter, except for those shown in Figures 7 and 8B. For these experiments, this potentiostat was used with a sampling interval of 0.1 seconds and a 3.2 Hz filter, or a 0.032 Hz filter, as indicated. This signal was recorded using the software provided by the manufacturer. Measurements of peak area, peak height, and amperometric current on the resulting current versus time curves were made using Graphpad Prism 6 software. All experiments were performed in 100 mM PBS buffer (pH = 7.4, 100 mM NaCl) at 33°C.

[0179] As disclosed in this document and shown throughout, the accumulation detection mode according to the embodiments of the present invention can be used to obtain superior detection compared to amperometry at low analyte concentrations. Petition 870250049834, dated 06 / 13 / 2025, pp. 84 / 97 76 / 76

[0180] Although the present invention has been illustrated and described with reference to certain exemplary embodiments, those generally skilled in the art will understand that various modifications and changes can be made to the described embodiments without departing from the character and scope of the present invention, as defined in the following claims. Petition 870250049834, dated 06 / 13 / 2025, pp. 85 / 97

Claims

1 / 5 CLAIMS 1. A method for detecting an analyte using a sensor, the sensor including an operating electrode, CHARACTERIZED in that the method comprises: providing the operating electrode with a detection element comprising an analyte-specific enzyme, a redox mediator and carbon nanotubes; providing the operating electrode to the analyte; accumulating charge on the redox mediator allowing the analyte to react with the analyte-specific enzyme for a stipulated period of time; connecting the operating electrode to a circuit after the stipulated period of time; and measuring a signal of the accumulated charge, wherein the redox mediator comprises a redox species selected from osmium, ruthenium, iron, cobalt and compounds or complexes thereof, coupled to a polymer selected from poly(vinylpyridine), poly(thiophene), poly(aniline), poly(pyrrole) and poly(acetylene).

2. Method according to claim 1, CHARACTERIZED in that the method further comprises connecting the working electrode to the circuit before supplying the working electrode to the analyte, and disconnecting the working electrode from the circuit after supplying the working electrode to the analyte.

3. Method according to claim 1, CHARACTERIZED in that after supplying the working electrode to the analyte, the method further comprises connecting the working electrode to the circuit, measuring an amperometric current and then disconnecting the working electrode from the circuit to begin accumulating charge.

4. Method according to claim 1, CHARACTERIZED in that the sensor is an enzymatic electrochemical biosensor.

5. Method, according to claim 1, CHARACTERIZED by the fact that the redox mediator is an immobilized redox polymer.

6. Method according to claim 1, CHARACTERIZED in that the analyte is selected from the group consisting of cortisol, glucose, lactate, 3-hydroxybutyrate, alcohol, pyruvate, glutamate, theophylline and creatinine.

7. Method according to claim 1, CHARACTERIZED in that the analyte-specific enzyme is selected from the group consisting of a nicotinamide adenine dinucleotide (NAD)-dependent dehydrogenase, a flavin adenine dinucleotide (FAD)-dependent oxidase, and a flavin mononucleotide (FMN)-dependent oxidase.

8. Method according to claim 1, CHARACTERIZED in that the analyte-specific enzyme is selected from the group consisting of 11β-hydroxy steroid dehydrogenase type 2 (1ie—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.

9. Method according to claim 1, CHARACTERIZED in that the analyte is at a concentration equal to or greater than 4.7 nanomolar.

10. Method, according to claim 1, CHARACTERIZED in that the measurement of the accumulated charge signal comprises measuring a peak height of the signal and / or measuring a peak area of ​​the signal.

11. Method according to claim 10, CHARACTERIZED in that it further comprises calibrating the measured peak height to obtain an analyte concentration.

12. Method, according to claim 10, CHARACTERIZED in that it further comprises calibrating the measured peak area to obtain an analyte concentration.

13. Method, according to claim 1, CHARACTERIZED by the fact that Petition 870260076213, dated 07 / 30 / 2026, page 17 / 20 3 / 5, the measurement of the accumulated charge signal comprises 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).

14. Method for continuously detecting an analyte in vivo using a sensor, the sensor comprising an operating electrode comprising an analyte-specific enzyme and a redox mediator, CHARACTERIZED in that the method comprises: providing the operating electrode to the analyte; accumulating charge derived from the analyte reacting with the analyte-specific enzyme and the redox mediator; and measuring a signal of the accumulated charge by measuring a peak height of the signal and / or measuring a peak area of ​​the signal.

15. Method according to claim 1, CHARACTERIZED in that the measurement of the accumulated charge signal comprises filtering the signal at a frequency of 3.2 hertz (Hz).

16. System for the detection of an analyte, CHARACTERIZED in that the system comprises: an operating electrode; a detection element disposed on the operating electrode, the detection element comprising an analyte-specific enzyme, a redox mediator and carbon nanotubes, the detection element configured to accumulate charge on the redox mediator allowing the analyte to react with the analyte-specific enzyme for a stipulated period of time; and a circuit configured to connect with the operating electrode after the stipulated period of time and to measure a signal of the accumulated charge, wherein the redox mediator comprises a redox species selected from osmium, ruthenium, iron, cobalt and compounds or complexes thereof, coupled to a Petition 870260076213, dated 07 / 30 / 2026, page.18 / 20 4 / 5 polymer selected from poly(vinylpyridine), poly(thiophene), poly(aniline), poly(pyrrole) and poly(acetylene), and wherein the measurement of the accumulated charge signal comprises filtering the signal at a frequency of 3.2 hertz (Hz).

17. System according to claim 16, CHARACTERIZED in that it further comprises an outer membrane covering at least the detection element.

18. System according to claim 16, CHARACTERIZED in that the analyte-specific enzyme is selected from the group consisting of a nicotinamide adenine dinucleotide (NAD)-dependent dehydrogenase, a flavin adenine dinucleotide (FAD)-dependent oxidase, and a flavin mononucleotide (FMN)-dependent oxidase.

19. System according to claim 16, CHARACTERIZED in that the analyte-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.

20. Method for continuously detecting an analyte in vivo using a sensor, the sensor including an operating electrode, CHARACTERIZED in that the method comprises: providing the operating electrode with an analyte-specific enzyme and a redox mediator; providing the operating electrode with the analyte; accumulating charge derived from the analyte reacting with the analyte-specific enzyme and the redox mediator for a stipulated period of time; connecting the operating electrode to a circuit after the stipulated period of time; and measuring a signal of the accumulated charge, wherein the measurement of the accumulated charge signal comprises recording the signal at a frequency of 3.2 hertz (Hz).

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