Thermally stable glucose limiting membrane for glucose sensors

Abstract

Embodiments of the present invention provide compositions useful in analyte sensors and methods for making and using such compositions and sensors. In typical embodiments of the present invention, the sensor is a glucose sensor that includes an analyte modulating membrane formed from a polymerization reaction mixture formed to include a limiting amount of catalyst and / or polycarbonate compound to provide improved material properties to such membranes, such as enhanced thermal stability and hydrolytic stability.

Description

[0001] This application is a divisional application of the invention patent application with application number 201980030243.0, application date May 16, 2019, entitled "Thermostabilized Glucose Restriction Membrane for Glucose Sensor".

[0002] Cross-referencing related applications This application claims priority to U.S. Patent Application Serial No. 15 / 981,681, entitled “THERMALLY STABLE GLUCOSE LIMITING MEMBRANE FOR GLUCOSE SENSORS,” filed May 16, 2018, pursuant to Section 120, the contents of which are incorporated herein by reference. Technical Field

[0003] This invention relates to biosensors such as glucose sensors for diabetes management and materials for manufacturing such sensors, such as polymeric compositions that can be used in biosensor membranes. Background Technology

[0004] Analytical sensors, such as biosensors, involve using biological elements to convert chemical analytes in a matrix into detectable signals. Many types of biosensors exist for detecting a wide variety of analytes. Perhaps the most studied type of biosensor is the amperometric glucose sensor, a device commonly used to monitor glucose levels in individuals with diabetes.

[0005] A typical glucose sensor operates based on the following chemical reaction:

[0006] Glucose oxidase catalyzes the reaction between glucose and oxygen to produce gluconic acid and hydrogen peroxide, as shown in Equation 1. H₂O₂ undergoes an electrochemical reaction, as shown in Equation 2, and the current is measured by a potentiostat. The stoichiometry of the reaction presents a challenge for developing sensors in vivo. Specifically, for optimal sensor performance, the sensor signal output should be determined solely by the analyte of interest (glucose), rather than by any co-matrix (O₂) or kinetic control parameters such as diffusion. If oxygen and glucose are present in equimolar concentrations, H₂O₂ is stoichiometrically related to the amount of glucose reacting at the enzyme; and the associated current generating the sensor signal is proportional to the amount of glucose reacting with the enzyme. However, if oxygen is insufficient to react all glucose with the enzyme, the current will be proportional to the oxygen concentration rather than the glucose concentration. Therefore, for the sensor to provide a signal dependent solely on glucose concentration, glucose must be the limiting agent, i.e., the O₂ concentration must exceed all potential glucose concentrations. However, a problem with using such glucose sensors in vivo is that the oxygen concentration relative to glucose is low when the sensor is implanted, which can impair the accuracy of sensor readings.

[0007] Numerous methods exist for addressing hypoxia problems. One approach involves using homogeneous polymer membranes with hydrophobic and hydrophilic regions that control oxygen and glucose permeability. For example, Van Antwerp et al. developed linear polyurea membranes comprising a combination of hydrophobic components of polyethylene glycol and silicone that allow high oxygen permeability and a hydrophilic component that allows limited glucose permeability (see, for example, U.S. Patents 5,777,060, 5,882,494, and 6,642,015). While possessing many useful and desirable properties, such polymeric compositions may exhibit some degradation over time under high temperature and high humidity conditions. Therefore, there is a need in the art for more robust polymeric membrane compositions that can, for example, address the hypoxia problem observed in glucose sensors incorporating glucose oxidase. Summary of the Invention

[0008] Embodiments of the present invention provide compositions that can be used in analyte sensors, as well as methods for manufacturing and using such compositions and sensors. In a typical embodiment of the invention, the sensor is a glucose sensor comprising an analyte-modulated membrane formed from a polymerization reaction mixture containing a limited amount of catalyst to provide such membrane with improved material properties, such as enhanced thermal and hydrolytic stability. As disclosed herein, when these polymer compositions are used to form an analyte-limiting membrane in a glucose sensor, the resulting sensor exhibits an enhanced long-term stability profile compared to conventional polymer compositions formed from reaction mixtures using conventional amounts of catalyst.

[0009] The invention disclosed herein has many embodiments. One embodiment of the invention is a method for increasing the thermal stability of a biocompatible membrane formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; a siloxane having an amino, hydroxyl, or carboxylic acid functional group at its terminal; and a catalyst. In this method, the reaction mixture is formed such that the catalyst is present in the reaction mixture in an amount less than 0.2% (e.g., 0.1%) of the reaction mixture components, thereby increasing the thermal stability of the biocompatible membrane compared to a comparable membrane formed from a reaction mixture in which the catalyst is present in a formulation in an amount greater than or equal to 0.2% of the reaction mixture components. Optionally, the reaction mixture further comprises additional components, such as polycarbonate diol.

[0010] Another embodiment of the present invention is an amperometric detection analyte sensor, the amperometric detection analyte sensor comprising: a substrate layer; a conductive layer disposed on the substrate layer and including a working electrode; an analyte sensing layer disposed on the conductive layer; and an analyte modulation layer disposed on the analyte sensing layer. In this embodiment, the analyte modulation layer is formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; a siloxane having an amino, hydroxyl, or carboxylic acid functional group at its terminal; and a catalyst. In this embodiment, the catalyst is present in the reaction mixture in an amount less than 0.2% of the reaction mixture components, such that the analyte modulation layer exhibits greater thermal stability compared to a comparable analyte modulation layer formed from a reaction mixture in which the catalyst is present in the formulation in an amount greater than or equal to 0.2% of the reaction mixture.

[0011] Another embodiment of the present invention is a method for manufacturing an analyte sensor for implantation in mammals. This method embodiment includes the steps of: providing a substrate layer; forming a conductive layer on the substrate layer, wherein the conductive layer comprises a working electrode; forming an analyte sensing layer on the conductive layer, wherein the analyte sensing layer comprises an oxidoreductase; and then forming an analyte modulation layer on the analyte sensing layer. In this embodiment, the analyte modulation layer is formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; a siloxane having an amino, hydroxyl, or carboxylic acid functional group at its terminal; and a catalyst present in the reaction mixture in an amount less than 0.2% (e.g., 0.1%) of the reaction mixture components, such that the reaction mixture exhibits greater thermal stability compared to a comparable analyte modulation layer formed from a reaction mixture in which the catalyst is present in a formulation in an amount greater than or equal to 0.2% of the reaction mixture components. Optionally, the reaction mixture further comprises additional components, such as polycarbonate diol.

[0012] Other objects, features, and advantages of the present invention will become apparent to those skilled in the art from the following detailed description. However, it should be understood that while some embodiments of the invention are indicated, the detailed description and specific examples are given by way of illustration rather than limitation. Many changes and modifications can be made within the scope of the invention without departing from its spirit, and the invention encompasses all such modifications. Attached Figure Description

[0013] Referring now to the accompanying drawings, in which similar reference numerals refer to corresponding parts throughout the following drawings: Figure 1 A schematic view of one embodiment of an amperometric analyte sensor having multiple layered materials / elements is provided according to one or more embodiments of the present invention.

[0014] Figure 2A -C illustrates the chemical structure of the raw material used in polycarbonate urea glucose restriction membrane (GLM) according to one or more embodiments of the present invention. Figure 2A The chemical structures of PDMS and Jeffamine are shown. Figure 2B 4,4'-methylenebis(cyclohexyl isocyanate) or HMDI and 4,4'-methylenebis(phenyl isocyanate) or MDI are shown. Figure 2C The chemical structure of polycarbonate diol is shown.

[0015] Figure 3 The GLM synthesis reaction according to one or more embodiments of the present invention is illustrated.

[0016] Figures 4A-4B A comparison of the morphology of various counter electrodes with that of the working electrode after testing according to one or more embodiments of the present invention is shown. Figure 4A The image shows bubbles (or pits) formed at the counter electrode after use. Bubble formation at the counter electrode may trigger stratification or unwanted biological responses (due to changes in texture or rough surfaces). Figure 4B It is shown that MDI_polycarbonate_GLM can enhance the adhesion of GLM, so that no bubbles (or pits) are generated at the counter electrode after use.

[0017] Figure 5 An in vitro SITS data comparison between a standard 2xGLM and a PCU_GLM (polycarbonate urea glucose restriction membrane) according to one or more embodiments of the present invention is shown.

[0018] Figure 6 The E3 sensor morphology is shown after 7-day SITS testing of a standard 2xGLM coated sensor and a PCU_GLM coated sensor according to one or more embodiments of the present invention.

[0019] Figures 7A-7B Data graphs from studies on the thermal degradation of various formulations and compositions of formulations are provided. Figure 7A Results of a thermal degradation study (as observed by molecular weight reduction) of polymeric materials suitable for use as biocompatible membranes (e.g., analyte modulation layers in glucose sensors) are provided at 45°C. Figure 7B Results from a similar thermal degradation study conducted at 60°C are provided. Figure 7A and Figure 7B In this formulation, the “control” material is formed from the polymerization reaction mixture, wherein the catalyst is present in the formulation at an amount greater than or equal to 0.2% of the polymerization reaction mixture, and the “new” material is formed from the polymerization reaction mixture, wherein the catalyst is present in the formulation at an amount less than 0.2% of the reaction mixture (in this case, 0.1%).

[0020] Figures 8A-8E Results of thermal degradation studies of various formulations and compositions thereof are shown according to one or more embodiments of the present invention. Figure 8A Results from thermal degradation studies are shown. Figures 8B-8D Compositions of various formulations are shown. Figure 8E Data are provided showing that the glucose permeability (Pg) of polycarbonate_GLM did not decrease after baking.

[0021] Figure 9A summary of the results of thermal / hydrolysis studies from various sample formulations according to one or more embodiments of the present invention is presented. Thermal degradation test results demonstrate that MDI and polycarbonate chains can contribute to (slow down) the GLM degradation process. Detailed Implementation

[0022] Unless otherwise defined, all technical terms, symbols, and other scientific terms or proprietary terms used herein are intended to have the meaning commonly understood by one of ordinary skill in the art to which this invention pertains. In some instances, for clarity and / or ease of reference, terms having a commonly understood meaning are defined herein, and the inclusion of such definitions herein should not necessarily be construed as representing a material difference from the meaning commonly understood in the art. Many of the techniques and procedures described or referenced herein are well understood and commonly employed by those skilled in the art using conventional methods. Procedures involving the use of commercially available kits and reagents are generally performed according to the manufacturer's defined schemes and / or parameters, as needed, unless otherwise indicated. Many terms are defined below. All publications referenced herein are incorporated by reference to disclose and describe methods and / or materials associated with the cited publications. The publications cited herein are referenced for their disclosure prior to the filing date of this application. Nothing herein shall be construed as an admission that the inventor has no right to advance the publication date by virtue of an earlier priority date or priority invention date. Furthermore, the actual publication date may differ from the date shown and requires independent verification.

[0023] It is important to note that, unless the context clearly indicates otherwise, the singular forms “a / an” and “described” as used herein and in the appended claims include plural references. Thus, for example, a reference to “an oxidoreductase” includes various such oxidoreductases and their equivalents known to those skilled in the art. All numbers listed in the specification and related claims that refer to values ​​that can be numerically characterized by values ​​other than integers (e.g., “50 mol%”) should be understood to be modified by the term “about”.

[0024] The term "analyte" as used herein is a broad term and is used in its common sense, including but not limited to substances or chemical components in analyzable fluids such as biological fluids (e.g., blood, interstitial fluid, cerebrospinal fluid, lymph, or urine). Analytes may include naturally occurring substances, artificial substances, metabolites, and / or reaction products. In some embodiments, the analyte for measurement by the sensing area, device, and method is glucose. However, other analytes are also contemplated, including but not limited to lactic acid. In some embodiments, naturally occurring salts, sugars, proteins, fats, vitamins, and hormones in blood or interstitial fluid may constitute analytes. Analytes may be naturally occurring in biological fluids or endogenous; for example, metabolites, hormones, antigens, antibodies, etc. Alternatively, analytes may be introduced into the human body or be exogenous, for example, contrast agents for imaging, radioactive isotopes, chemical reagents, synthetic blood based on fluorocarbons, or drugs or pharmaceutical compositions, including but not limited to insulin. Metabolites of drugs and pharmaceutical compositions are also intended analytes.

[0025] As used herein, the term "sensor" is a broad term and is used in its common sense, encompassing, but not limited to, the detection of one or more analyte components by an analyte monitoring device. In one embodiment, the sensor comprises an electrochemical cell having a working electrode, a reference electrode, and optionally a counter electrode, the counter electrode passing through and fixed within the sensor body, thereby forming an electrochemically reactive surface at one location on the body, an electrical connection at another location on the body, and a membrane system adhered to the body and covering the electrochemically reactive surface. During normal operation of the sensor, a biological sample (e.g., blood or interstitial fluid) or a portion thereof comes into contact with an enzyme (e.g., glucose oxidase) (either directly or after passing through one or more membranes or domains); the reaction of the biological sample (or a portion thereof) results in the formation of a reaction product that allows for the determination of analyte levels in the biological sample.

[0026] As discussed in detail below, embodiments of the present invention relate to the use of electrochemical sensors exhibiting a range of novel material and functional elements. Such sensors incorporate novel polymeric compositions to form robust analyte-modulated membranes possessing a unique set of technically desirable material properties, such as increased thermal stability. The electrochemical sensors of the present invention are designed to measure an analyte of interest (e.g., glucose) or indicate the concentration of an analyte in a fluid or the concentration of a substance present. In some embodiments, the sensor is a continuous device, such as a subcutaneous, percutaneous, or intravascular device. In some embodiments, the device can analyze multiple intermittent blood samples. The sensor embodiments disclosed herein can use any known method (including invasive, minimally invasive, and non-invasive sensing techniques) to provide an output signal indicating the concentration of an analyte of interest. Typically, sensors are of the type that sense the product or reactant of an enzymatic reaction between an analyte and an enzyme in the presence of oxygen to serve as a measure of the analyte in vivo or in vitro. Such sensors comprise a polymeric membrane surrounding an enzyme through which the analyte migrates before reacting with the enzyme. The product is then measured using an electrochemical method, and thus the output of the electrode system serves as a measure of the analyte. In some embodiments, the sensor may use current titration, coulometric titration, conductivity titration, and / or potentiometric titration techniques to measure the analyte.

[0027] Analyte modulation compositions (such as analyte modulation compositions that can be used as glucose limiting membranes in amperometric glucose sensors) comprise polymeric compositions formed from biocompatible polyurea materials (see, for example, the case of biocompatible polyurea materials, incorporated herein by reference). Such compositions can exhibit stable glucose and oxygen permeability, low protein adsorption, and biocompatibility. However, due to the content of PEG chains, certain degradation issues arise under high temperature and / or high humidity conditions. As detailed below, we have discovered that certain carbonate and aromatic isocyanate compounds can be added to the polymerization reaction to replace some portions of the polymer chain elements of PDMS and HMDI. Two compounds have been found to increase the heat resistance and hydrolysis resistance of these polymers under high temperature and high humidity conditions. In addition, the chemical structure of the compounds provides evidence that such compositions have very good electron beam resistance. Carbonate materials that can be used in embodiments of the present invention comprise, but are not limited to, polycarbonate diols (e.g., butanediol or hexanediol or similar compounds). In illustrative embodiments of the present invention, the Mw of the carbonate material is 500 to 2000 Daltons. The aromatic isocyanate materials that can be used in the embodiments of the present invention include, but are not limited to, MDI or similar compounds.

[0028] In addition to limiting the amount of catalyst in the polymerization reaction mixture, the addition of MDI can improve the heat resistance and electron beam resistance of the polymeric composition used as an analyte modulation (e.g., glucose restriction) composition through its benzene ring structure. The benzene ring also acts as a good free radical scavenger to prevent oxidation of the polymeric components. Polycarbonate diols can provide better heat resistance and hydrolysis resistance through their carbonate structure (as opposed to ether or ester chains). Adding polycarbonate segments to the polymer backbone can prevent unwanted deformation of the polymeric composition layer disposed on the electrode of an amperometric glucose sensor. Both gases and water are generated on the counter electrode between the analyte sensing layer (e.g., an analyte sensing layer containing enzymes such as GOX) and the analyte molding (e.g., a glucose restriction membrane) layer, which can lead to sensor malfunction (signal drift) after prolonged use. In this case, the polycarbonate segments in the GLM backbone can prevent / reduce the chain rotation of PDMS in the GLM membrane, thus the glucose permeability (Pg) of the GLM does not gradually decrease over time because the hydrophilic chains (Jeffamine or PEG) are wrapped / encapsulated by the hydrophobic PDMS chains, especially for low-Pg GLMs. In some embodiments, to produce a uniform urethane / urea copolymer, the synthesis will involve three feedstock injections at different timings. The feedstocks are injected at a 4-2-4 ratio at times 0, 4, and 12 hours, respectively. The addition of polycarbonate chains to the GLM can prevent Pg changes / reductions due to the rotation / entanglement of PDMS chains over time, especially for low-Pg GLM membranes. To reduce thermal / radiative / oxidative degradation, the desired MDI content in the final polymer can be 2% to 25%. To prevent membrane deformation or Pg reduction due to the rotation of silicone chains over time, the desired polycarbonate content in the final polymer can be 8% to 30%. The polycarbonate GLM exhibits good adhesion to AP, with no additional pits (bubbles) formed after testing.

[0029] The embodiments of the invention disclosed herein provide, for example, sensor types for subcutaneous or transcutaneous monitoring of blood glucose levels in diabetic patients. Various implantable electrochemical biosensors have been developed to treat diabetes and other life-threatening diseases. Many existing sensor designs utilize some form of immobilized enzyme to achieve their biological specificity. The embodiments of the invention described herein can be adapted and implemented using a wide variety of known electrochemical sensors, including, for example, U.S. Patent Application No. 20050115832, U.S. Patent Nos. 6,001,067, 6,702,857, 6,212,416, 6,119,028, 6,400,974, 6,595,919, 6,141,573, 6,122,536, and 6,51... Nos. 2,939, 5,605,152, 4,431,004, 4,703,756, 6,514,718, 5,985,129, 5,390,691, 5,391,250, 5,482,473, 5,299,571, 5,568,806, 5,494,562, 6,120,676, 6,542,765 and PCT International Publication No. WO WO 01 / 58348, WO 04 / 021877, WO 03 / 034902, WO 03 / 035117, WO 03 / 035891, WO 03 / 023388, WO 03 / 022128, WO 03 / 022352, WO 03 / 023708, WO 03 / 036255, WO 03 / 036310, WO 08 / 042625 and WO 03 / 074107, and European Patent Application EP 1153571, the contents of each of these documents are incorporated herein by reference.

[0030] As discussed in detail below, embodiments of the invention disclosed herein provide sensor elements with enhanced material properties and / or architectural configurations, and sensor systems configured to include such elements (e.g., sensor systems including sensors and associated electronic components such as monitors, processors, etc.). This disclosure further provides methods for manufacturing and using such sensors and / or architectural configurations. While some embodiments of the invention relate to glucose sensors and / or lactate sensors, the various elements disclosed herein (e.g., analyte modulation membranes made from polycarbonate polymer compositions) are adaptable for use with any of the wide variety of sensors known in the art. Analyte sensor elements, architectures, and methods for manufacturing and using these elements disclosed herein can be used to create a variety of layered sensor structures. Such sensors of the present invention exhibit surprising flexibility and versatility, allowing a wide variety of sensor configurations to be designed to examine the properties of a wide variety of analytes.

[0031] Specific aspects of the embodiments of the present invention will be discussed in detail in the following sections.

[0032] Typical elements, configurations, and analyte sensors of the present invention The optimized sensor element of the present invention A wide variety of sensors and sensor elements are known in the art, including amperometric detection sensors for detecting and / or measuring bioanalytes such as glucose. Many glucose sensors are based on oxygen (Clark-type) amperometric detection transducers (see, for example, Yang et al., *Electroanalysis*, 1997, 9, 16: 1252-1256; Clark et al., *Acad. Sci.*, Ann. NY, 1962, 102, 29; Updike et al., *Nature*, 1967, 214, 986; and Wilkins et al., *Med. Engin. Physics*, 1996, 18, 273, 3-51). Many in vivo glucose sensors utilize hydrogen peroxide-based amperometric detection transducers because the transducers are relatively easy to manufacture and can be readily miniaturized using conventional techniques. However, one problem associated with the use of certain amperometric detection transducers involves suboptimal reaction stoichiometry. As discussed in detail below, these problems are addressed by using one or more polycarbonate polymer membranes disclosed herein, which can modulate the transport properties of different compounds whose reactions generate signals at an amperometric detection transducer based on hydrogen peroxide. Therefore, these membranes can be used with, for example, various H₂O₂-based analyte sensors that benefit from optimized reaction stoichiometry.

[0033] As described above, embodiments of the present invention comprise sensor membranes made in the form of reaction mixtures to contain limited amounts of a catalyst and / or a polycarbonate polymer composition. As known in the art, polymers comprise long or larger molecules consisting of chains or networks of a plurality of repeating units, which are formed by chemically bonding together a plurality of identical or similar small molecules called monomers. Copolymers or heteropolymers are polymers derived from two (or more) monomer species, as opposed to homopolymers using only one monomer. Copolymers can also be described in terms of the presence or arrangement of branches in the polymer structure. Linear copolymers consist of a single backbone, while branched copolymers consist of a single backbone and one or more polymer side chains. Sensor membranes made from the polycarbonate polymer compositions disclosed herein can optimize analyte sensor functionality, including sensor sensitivity, stability, and hydration profile. Furthermore, by optimizing the stoichiometry of reactant species over a sensor temperature range, the membranes disclosed herein can optimize the chemical reactions that produce key measurable signals associated with the level of the analyte of interest (e.g., glucose). The following sections describe illustrative sensor elements, sensor configurations, and method embodiments of the present invention.

[0034] The polymer materials disclosed herein can be used as biocompatible membranes in various situations, such as glucose limiting membranes (GLMs). However, polymers typically degrade over time due to oxidation, UV light, heat, hydrolysis, or other processes. In such cases, trace amounts of tin catalyst residues in the polymer mixtures used to form biocompatible membranes such as GLMs have been found to accelerate GLM degradation over time. For this reason, older sensors may perform slightly worse than newly manufactured sensors due to the gradual degradation of GLMs. While not bound by any specific scientific theory or mechanism of action, it is believed that trace amounts of tin catalyst residues in GLMs may further trigger immune responses and sensitivity loss. In this case, it has been further found that reducing the amount of tin catalyst used in the GLM synthesis process (e.g., by 50%) unexpectedly produces membranes with increased resistance to thermal degradation and improves the quality of GLMs (biocompatibility and thermal stability) as well as the in vivo performance of the glucose sensor. This greater GLM stability further extends the sensor's storage life and improves its biocompatibility without any significant changes in the manufacturing process.

[0035] The invention disclosed herein has many embodiments. One embodiment of the invention is a method for increasing the thermal stability of a biocompatible membrane formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; a siloxane having an amino, hydroxyl, or carboxylic acid functional group at its terminal; and a catalyst. In this method, the reaction mixture is formed such that the catalyst is present in the reaction mixture in an amount less than 0.2% (e.g., 0.1%) of the reaction mixture components, thereby increasing the thermal stability of the biocompatible membrane compared to a comparable membrane formed from a reaction mixture in which the catalyst is present in a formulation in an amount greater than or equal to 0.2% of the reaction mixture components. Optionally, the reaction mixture (e.g., at 60°C) uses an organic solvent such as tetrahydrofuran and further comprises additional components such as polycarbonate diol. The thermal stability of various biocompatible membranes manufactured in this manner can be measured by various practices accepted by existing technology, such as by observing the change in molecular weight of the biocompatible membrane during maintenance at 60°C for at least 3, 5, or 7 days (see, for example, Figure 7).

[0036] In some embodiments of the invention, a tin catalyst (e.g., dibutyltin bis(2-hexanoate)) is present in the reaction mixture in an amount (e.g., 0.1%) less than 0.19%, 0.17%, 0.15%, 0.13%, or 0.11% of the reaction mixture. In some embodiments of the invention, the diisocyanate comprises hexamethylene diisocyanate and / or methylene diphenyl diisocyanate, and / or the hydrophilic polymer comprising a hydrophilic diol or hydrophilic diamine comprises JEFFAMINE, and / or the siloxane having an amino, hydroxyl, or carboxylic acid functional group at the end comprises polydimethylsiloxane, and / or the polycarbonate diol comprises poly(1,6-hexyl carbonate) diol and / or poly(1,6-hexyl-1,5-pentyl carbonate) diol. For example, in some embodiments of the invention, the diisocyanate comprises 17% to 23% by weight of hexamethylene diisocyanate and 0% to 8.5% by weight of methylene diphenyl diisocyanate, and JEFFAMINE comprises 28% to 51% by weight of JEFFAMINE 600 and / or JEFFAMINE 900, and the polydimethylsiloxane comprises 14% to 48% by weight of polydimethylsiloxane-A15, and the polycarbonate diol comprises 7.5% to 19% by weight of poly(1,6-hexyl carbonate) diol. In specific illustrative embodiments, the diisocyanate comprises about 22% hexamethylene diisocyanate and about 3.5% methylene diphenyl diisocyanate, JEFFAMINE comprises about 45% JEFFAMINE 600 and / or JEFFAMINE 900, the polydimethylsiloxane comprises about 22.5% polydimethylsiloxane-A15, and the polycarbonate diol comprises about 7.5% poly(1,6-hexyl carbonate) diol. Optionally, in this method, water is added as a chain extender to the reaction mixture of the polyurea-polyurethane copolymer. The illustrative reaction mixture of the present invention is shown in Tables 1 and 2 below.

[0037] Table 1

[0038] Table 2

[0039] Another embodiment of the present invention is an amperometric detection analyte sensor, the amperometric detection analyte sensor comprising: a substrate layer; a conductive layer disposed on the substrate layer and including a working electrode; an analyte sensing layer disposed on the conductive layer; and an analyte modulation layer disposed on the analyte sensing layer. In this embodiment, the analyte modulation layer is formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; a siloxane having an amino, hydroxyl, or carboxylic acid functional group at its terminal; and a catalyst. In this embodiment, the catalyst is present in the reaction mixture in an amount less than 0.2% of the reaction mixture components, such that the analyte modulation layer exhibits greater thermal stability compared to a comparable analyte modulation layer formed from a reaction mixture in which the catalyst is present in the formulation in an amount greater than or equal to 0.2% of the reaction mixture. Optionally, the reaction mixture further comprises additional components, such as polycarbonate diol.

[0040] In a typical embodiment, the analyte sensor is an implantable glucose sensor. Optionally, the analyte sensor further includes at least one of the following: a protein layer disposed on the analyte sensing layer; or a capping layer disposed on the analyte sensor device, the capping layer including pores positioned on the capping layer to facilitate the analyte present in the in vivo environment from contacting and diffusing through the analyte modulation layer; and contacting the analyte sensing layer. In some of these analyte sensors, the conductive layer includes a plurality of electrodes comprising a working electrode, a counter electrode, and a reference electrode. For example, in one embodiment, the conductive layer includes a plurality of working electrodes and / or counter electrodes and / or reference electrodes; and optionally, the plurality of working electrodes, counter electrodes, and reference electrodes are grouped together as units and distributed in position on the conductive layer in a unit repeating pattern.

[0041] Another embodiment of the present invention is a method for manufacturing an analyte sensor for implantation in mammals. This method embodiment includes the steps of: providing a substrate layer; forming a conductive layer on the substrate layer, wherein the conductive layer comprises a working electrode; forming an analyte sensing layer on the conductive layer, wherein the analyte sensing layer comprises an oxidoreductase; and then forming an analyte modulation layer on the analyte sensing layer. In this embodiment, the analyte modulation layer is formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; a siloxane having an amino, hydroxyl, or carboxylic acid functional group at its terminal; and a catalyst present in the reaction mixture in an amount less than 0.2% (e.g., 0.1%) of the reaction mixture components, such that the reaction mixture exhibits greater thermal stability compared to a comparable analyte modulation layer formed from a reaction mixture in which the catalyst is present in a formulation in an amount greater than or equal to 0.2% of the reaction mixture components. Optionally, the reaction mixture further comprises additional components, such as polycarbonate diol.

[0042] In one method of manufacturing an analyte sensor for implantation in mammals, the diisocyanate comprises hexamethylene diisocyanate and / or methylene diphenyl diisocyanate, JEFFAMINE comprises about 45% JEFFAMINE 600 and / or JEFFAMINE 900, the polydimethylsiloxane comprises about 22.5% polydimethylsiloxane-A15, and the polycarbonate diol comprises about 7.5% poly(1,6-hexyl carbonate) diol. Typically, in this embodiment, the catalyst (e.g., dibutyltin bis(2-hexanoate)) is present in the reaction mixture in an amount less than 0.19%, 0.17%, 0.15%, 0.13%, or 0.11% of the reaction mixture (e.g., about 0.1%).

[0043] An amperometric detection sensor design used in conjunction with embodiments of the present invention includes multiple layered elements, including, for example, a substrate layer having electrodes, an analyte sensing layer (e.g., an analyte sensing layer including glucose oxidase), and an analyte modulation layer that functions in analyte diffusion control (e.g., to modulate the amount of glucose and oxygen exposed to the analyte sensing layer). Figure 1An example of such a sensor is illustrated herein. The layered sensor design, incorporating the polycarbonate polymeric composition disclosed herein as an analyte modulation layer, exhibits a range of material properties that overcome challenges observed in various sensors, including in vivo implanted electrochemical glucose sensors. For example, sensors designed to measure analytes in aqueous environments (e.g., in vivo implanted sensors) typically require wetting of the layers before and during the measurement of accurate analyte readings. Because the properties of the materials can affect their hydration rate, the material properties of the membrane used in an aqueous environment ideally promote sensor wetting, for example, enabling the measurement of the time between the sensor's introduction into the aqueous environment and its ability to provide an accurate signal corresponding to the analyte concentration in said environment. Embodiments of the invention, including polycarbonate polymeric compositions, address these problems by promoting sensor hydration.

[0044] Furthermore, in the case of electrochemical glucose sensors that utilize the chemical reaction between glucose and glucose oxidase to generate a measurable signal, the material of the analyte modulation layer should not exacerbate (and ideally should reduce) the condition known in the art as the "hypoxia problem." Specifically, because glucose oxidase-based sensors require both oxygen (O2) and glucose to generate a signal, an excess of oxygen relative to an excess of glucose is necessary for the operation of glucose oxidase-based glucose sensors. However, since the oxygen concentration in subcutaneous tissue is much lower than the glucose concentration, oxygen may be the limiting reactant in the reaction between glucose, oxygen, and glucose oxidase in the sensor, thus impairing the sensor's ability to generate a signal that is strictly dependent on glucose concentration. In this case, because the properties of the material may affect the rate at which the compound diffuses through the material to the site of the measurable chemical reaction, the material properties of the analyte modulation layer used in electrochemical glucose sensors that utilize the chemical reaction between glucose and glucose oxidase to generate a measurable signal should not, for example, be more favorable to glucose diffusion than oxygen in a way that causes the hypoxia problem. The embodiments of the invention, including the polycarbonate polymer compositions disclosed herein, do not cause the hypoxia problem but rather mitigate it.

[0045] Furthermore, sensor designs using the polycarbonate polymer compositions disclosed herein as analyte modulation layers can overcome the complexities observed when using sensor materials that exhibit different diffusion profiles (e.g., the rate at which analytes diffuse through the sensor material) at different temperatures. Specifically, to achieve optimized sensor performance, the sensor signal output over a given temperature range should be determined solely by the level of the analyte of interest (e.g., glucose), rather than by any co-matrix (e.g., O2) or kinetic control parameters. However, as is known in the art, the diffusion of compounds through a polymer matrix can be temperature-dependent. Such a temperature-dependent diffusion profile can affect the stoichiometry of the reaction upon which the generation of the sensor signal depends, in cases where an analyte (e.g., glucose) diffuses through the polymer to react at a site with another compound (e.g., glucose oxidase), thus hindering the effort of those skilled in the art to make the sensor signal output depend solely on the concentration of the analyte of interest within a given temperature range. Therefore, analyte modulation compositions made from materials having a stable analyte (e.g., glucose) diffusion profile over a given temperature range (e.g., 22 to 40 degrees Celsius) address this problem.

[0046] The invention disclosed herein provides polycarbonate polymeric compositions that can be used as membranes for biosensors, such as amperometric glucose sensors. Embodiments of the invention include, for example, sensors having multiple layered elements comprising an analyte-limiting membrane comprising the polycarbonate polymeric composition. Such polymeric membranes can specifically be used to construct electrochemical sensors for in vivo use. The membrane embodiments of the invention allow for a combination of desired properties, including enhanced hydration profile and stable permeability to molecules such as glucose over a certain temperature range. Additionally, these polymeric membranes exhibit good mechanical properties for use as external polymeric membranes. Therefore, glucose sensors incorporating such polymeric membranes exhibit a highly desirable in vivo performance profile.

[0047] Embodiments of the present invention include both materials (e.g., polycarbonate polymer compositions) and architectures designed to enhance sensor performance. For example, in some embodiments of the invention, the conductive layer includes a plurality of working electrodes and / or counter electrodes and / or reference electrodes (e.g., three working electrodes, one reference electrode, and one counter electrode) to, for example, avoid problems associated with poor sensor hydration and / or provide redundant sensing capabilities. Optionally, the plurality of working electrodes, counter electrodes, and reference electrodes are configured together as units and distributed positionally on the conductive layer in a repeating unit pattern. In some embodiments of the invention, the base layer is made of a flexible material that allows the sensor to twist and bend when implanted in vivo; and the electrodes are grouped in a configuration that facilitates in vivo fluid contact with at least one of the working electrodes when twisting and bending occurs during implantation of the sensor device in vivo. In some embodiments, the electrodes are grouped in a configuration that allows the sensor to continue operating if a portion of the sensor having one or more electrodes is removed from the in vivo environment and exposed to an ex vivo environment. Typically, a sensor is operatively coupled to: a sensor input capable of receiving signals from the sensor based on sensed analytes; and a processor coupled to the sensor input, wherein the processor is capable of characterizing one or more signals received from the sensor. In some embodiments of the invention, pulsed voltages are used to obtain signals from one or more electrodes of the sensor.

[0048] The sensors disclosed herein can be made from a wide variety of materials known in the art. In one illustrative embodiment of the invention, the analyte modulation layer comprises a polyurethane / polyurea polymer formed from a mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; and a siloxane having amino, hydroxyl, or carboxylic acid functional groups at its ends; using this polymer, the polycarbonate of the branched acrylate polymer formed from the mixture comprises: butyl acrylate, propyl acrylate, ethyl acrylate, or methyl acrylate; amino acrylate; siloxane-acrylate; and poly(ethylene oxide)-acrylate. Optionally, additional materials may be included in these polymer blends. For example, some embodiments of the branched acrylate polymer are formed from a reaction mixture comprising a hydroxy acrylate compound (e.g., 2-hydroxyethyl methacrylate).

[0049] As used herein, the term "polyurethane / polyurea polymer" refers to a polymer containing urethane bonds, urea bonds, or a combination thereof. As is known in the art, polyurethane is a polymer composed of organic unit chains linked by urethane (carbamate) chains. Polyurethane polymers are typically formed by stepwise growth polymerization in the presence of a catalyst, reacting a monomer containing at least two isocyanate functional groups with another monomer containing at least two hydroxyl groups (alcohols). Polyurea polymers are derived from the reaction product of an isocyanate component and a diamine. Typically, such polymers are formed by combining a diisocyanate with an alcohol and / or an amine. For example, combining isophorone diisocyanate with PEG 600 and aminopropyl polysiloxane under polymerization conditions provides a polyurethane / polyurea composition having both urethane (carbamate) chains and urea bonds. Such polymers are known in the art and described, for example, in the following documents: U.S. Patents 5,777,060, 5,882,494, and 6,632,015; and PCT Publications WO 96 / 30431; WO 96 / 18115; WO 98 / 13685; and WO 98 / 17995, the contents of each of which are incorporated herein by reference.

[0050] The polyurethane / polyurea compositions of the present invention are prepared from bioacceptable polymers whose hydrophobic / hydrophilic balance can be varied over a wide range to control the ratio of oxygen diffusion coefficient to glucose diffusion coefficient and to match this ratio to the design requirements of electrochemical glucose sensors intended for in vivo use. Such compositions can be prepared by conventional methods of polymerization of the aforementioned monomers and polymers. The resulting polymers are soluble in solvents such as acetone or ethanol and can be formed into films from solutions by dip coating, spray coating, or spin coating.

[0051] The diisocyanate used in this embodiment of the invention is a diisocyanate commonly used in the preparation of biocompatible polyurethanes. Such diisocyanates are described in detail in *Szycher, Seminar on Advances in Medical Grade Polyurethanes*, Technomic Publishing, (1995) and include both aromatic and aliphatic diisocyanates. Examples of suitable aromatic diisocyanates include toluene diisocyanate, 4,4'-methylene diphenyl diisocyanate, 3,3'-dimethyl-4,4'-biphenyl diisocyanate, naphthalene diisocyanate, and terephthalic diisocyanate. Suitable aliphatic diisocyanates include, for example, 1,6-hexamethylene diisocyanate (HDI), trimethylhexyl diisocyanate (TMDI), trans-1,4-cyclohexane diisocyanate (CHDI), 1,4-cyclohexanebis(methylene isocyanate) (BDI), 1,3-cyclohexanebis(methylene isocyanate) (H6 XDI), isophorone diisocyanate (IPDI), and 4,4'-methylenebis(cyclohexyl isocyanate) (H2MDI). In some embodiments, the diisocyanate is isophorone diisocyanate, 1,6-hexamethylene diisocyanate, or 4,4'-methylenebis(cyclohexyl isocyanate). Many of these diisocyanates are available from commercial sources such as Aldrich Chemical Company (Milwaukee, Wisconsin, USA) or can be readily prepared using standard synthetic methods using documented procedures.

[0052] The amount of diisocyanate in the reaction mixture used in the polyurethane / polyurea polymer composition is typically about 50 mol% relative to the combination of the remaining reactants. More specifically, the amount of diisocyanate used to prepare the polyurethane / polyurea polymer will be sufficient to provide at least about 100% of the -NCO groups required for reaction with the hydroxyl or amino groups of the remaining reactants. For example, a polymer prepared using x moles of diisocyanate will use a moles of a hydrophilic polymer (diol, diamine, or a combination thereof), b moles of a silicone polymer with functionalized ends, and c moles of chain extender, such that x = a + b + c, where it should be understood that c can be zero.

[0053] Another reactant used to prepare the polyurethane / polyurea polymer described herein is a hydrophilic polymer. The hydrophilic polymer may be a hydrophilic diol, a hydrophilic diamine, or a combination thereof. The hydrophilic diol may be a poly(alkylene)ethylene glycol, a polyester-based polyol, or a polycarbonate polyol. As used herein, the term "poly(alkylene)ethylene glycol" refers to a polymer of a lower alkylene glycol such as polyethylene glycol, polyethylene glycol, and polytetramethylene ether glycol (PTMEG). The term "polyester-based polyol" refers to a polymer in which the R group is a lower alkylene group, such as ethylene, 1,3-propylene, 1,2-propylene, 1,4-butene, 2,2-dimethyl-1,3-propylene, etc. (e.g., as depicted in Figure 4 of U.S. Patent No. 5,777,060). Those skilled in the art will also understand that the diester portion of the polymer may also differ from the hexadecanoic acid shown. For example, while Figure 4 of U.S. Patent No. 5,777,060 illustrates the adipic acid component, the present invention also considers the use of succinate, glutarate, etc. The term "polycarbonate polyol" refers to a polymer having hydroxyl functionality at the chain ends and carbonate functionality within the polymer chain. The alkyl portion of the polymer is typically composed of C2 to C4 aliphatic groups, or in some embodiments, of longer-chain aliphatic, alicyclic, or aromatic groups. The term "hydrophilic diamine" refers to any of the above-described hydrophilic diols, wherein the terminal hydroxyl group has been substituted with a reactive amino group or wherein the terminal hydroxyl group has been derived to produce a chain extension having a terminal amino group. For example, a hydrophilic diamine is "diaminopoly(oxyethylene)", which is a poly(alkylene) glycol wherein the terminal hydroxyl group is replaced by an amino group. The term "diaminopoly(oxyethylene)" also refers to a poly(alkylene) glycol having an amino alkyl ether group at the chain ends. A suitable example of a diaminopoly(oxyolefin) is poly(propylene glycol) bis(2-aminopropyl ether). Many of the above polymers are available from Aldridge Chemical Company. Alternatively, conventional methods known in the art can be used for the synthesis of the polymers.

[0054] The amount of hydrophilic polymer used to prepare the linear polymer composition is typically from about 10 mol% to about 80 mol% relative to the diisocyanate used. Typically, this amount is from about 20 mol% to about 60 mol% relative to the diisocyanate. When using lower amounts of hydrophilic polymer, it is common to include a chain extender.

[0055] The silicone-containing polyurethane / polyurea polymers used in this invention are generally linear, exhibit excellent oxygen permeability, and are substantially free of glucose permeability. Typically, the silicone polymer is a polydimethylsiloxane having two reactive functional groups (i.e., a functionality of 2). The functional groups can be, for example, hydroxyl, amino, or carboxylic acid groups, but are typically hydroxyl or amino. In some embodiments, a combination of silicone polymers can be used, wherein the first portion comprises a hydroxyl group and the second portion comprises an amino group. Typically, the functional groups are located at the chain ends of the silicone polymer. Many suitable silicone polymers are commercially available from sources such as The Dow Chemical Company (Midland, Michigan, USA) and General Electric Company (Silicones Division, Schenectady, NY, USA). Other silicone polymers can be prepared from commercially available siloxanes (United Chemical Technologies, Bristol, Pa., USA) using general synthetic methods known in the art (see, for example, U.S. Patent No. 5,777,060). For use in this invention, the silicone polymer is typically a silicone polymer with a molecular weight of about 400 to about 10,000, and more typically a silicone polymer with a molecular weight of about 2,000 to about 4,000. The amount of silicone polymer incorporated into the reaction mixture will depend on the desired properties of the resulting polymer forming a biocompatible membrane. For compositions in which lower glucose permeability is desired, a larger amount of silicone polymer may be used. Alternatively, for compositions in which higher glucose permeability is desired, a smaller amount of silicone polymer may be used. Typically, for glucose sensors, the amount of silicone polymer is 10 mol% to 90 mol% relative to diisocyanate. Typically, the amount is about 20 mol% to 60 mol% relative to diisocyanate.

[0056] In one set of embodiments, the reaction mixture for preparing the biocompatible membrane will also contain a chain extender, which is an aliphatic or aromatic diol, an aliphatic or aromatic diamine, an alkanolamine, or a combination thereof (e.g., as depicted in Figure 8 of U.S. Patent No. 5,777,060). Examples of suitable aliphatic chain extenders include ethylene glycol, propylene glycol, 1,4-butanediol, 1,6-hexanediol, ethanolamine, ethylenediamine, butanediamine, and 1,4-cyclohexanediethanol. Aromatic chain extenders include, for example, p-bis(2-hydroxyethoxy)benzene, m-bis(2-hydroxyethoxy)benzene, Ethacure 100 (a mixture of two isomers of 2,4-diamino-3,5-diethyltoluene), Ethacure 300 (2,4-diamino-3,5-di(methylthio)toluene), 3,3'-dichloro-4,4'-diaminodiphenylmethane, Polacure (740M) (trimethylene glycol bis(p-aminobenzoic acid) ester), and diphenylaminomethane. The incorporation of one or more of these chain extenders typically provides additional physical strength to the resulting biocompatible membrane without significantly increasing the polymer's glucose permeability. Chain extenders are typically used when employing low amounts (i.e., 10-40 mol%) of hydrophilic polymers. Specifically, in some compositions, the chain extender is diethylene glycol present in a form of about 40 mol% to 60 mol% relative to the diisocyanate.

[0057] The polymerization of the above reactants can be carried out in batch form or in a solvent system. The use of a catalyst is partially necessary but not essential. Suitable catalysts include dibutyltin bis(2-hexanoate) (CAS No.: 2781-10-4), dibutyltin diacetate, triethylamine, and combinations thereof. Typically, dibutyltin bis(2-hexanoate) is used as the catalyst. The typical amount of this catalyst used in formulations is 0.05% to 0.2% (w / w ratio). Bulk polymerization is typically carried out at an initial temperature of about 25°C (ambient temperature) to about 50°C to ensure thorough mixing of the reactants. Exothermic reactions are typically observed during mixing of the reactants, with the temperature rising to about 90-120°C. After the initial exothermic reaction, the reaction flask can be heated to 75°C to 125°C, with 90°C to 100°C being an exemplary temperature range. Heating is typically performed for one to two hours. Solution polymerization can be carried out in a similar manner. Solvents suitable for solution polymerization include dimethylformamide, dimethyl sulfoxide, dimethylacetamide, halogenated solvents such as 1,2,3-trichloropropane, and ketones such as 4-methyl-2-pentanone. THF is typically used as the solvent. When polymerization is carried out in a solvent, heating of the reaction mixture is usually performed for three to four hours.

[0058] Polymers prepared by bulk polymerization typically dissolve in dimethylformamide and precipitate out of water. Polymers prepared in solvents immiscible with water can be isolated by vacuum stripping of the solvent. These polymers then dissolve in dimethylformamide and precipitate out of water. After thorough washing with water, the polymers can be vacuum dried to constant weight at approximately 50°C.

[0059] The membrane can be prepared by dissolving a dried polymer in a suitable solvent and casting the membrane onto a glass plate. The choice of a suitable solvent for casting typically depends on the specific polymer and the volatility of the solvent. Typically, solvents are THF, CHCl3, CH2Cl2, DMF, IPA, or combinations thereof. More typically, the solvent is THF or DMF / CH2Cl2 (2 / 98 wt%). The solvent is removed from the membrane, the resulting membrane is fully hydrated, its thickness is measured, and the water pickup rate is determined. Membranes used in this invention typically have a water pickup rate of about 20 to about 100 wt%, typically 30 to about 90 wt%, and more typically 40 to about 80 wt%.

[0060] For both individual polymer compositions and the polycarbonate polymeric membranes of the present invention, the oxygen diffusion coefficient and glucose diffusion coefficient can also be determined. Methods for determining the diffusion coefficients are known to those skilled in the art, and examples are provided below. Some embodiments of the biocompatible membranes described herein typically have a diffusion coefficient of about 0.1 x 10⁻⁶. -6 cm 2 / sec to approximately 2.0 x 10 -6 cm 2 Oxygen diffusion coefficient per second (D) 氧 ) and approximately 1 x 10 -9 cm 2 / sec to approximately 500 x 10 -9 cm 2 glucose diffusion coefficient per second (D) 葡萄糖 More typically, the glucose diffusion coefficient is approximately 10 x 10⁻⁶. -9 cm 2 / sec to approximately 200 x 10 -9 cm 2 / sec.

[0061] Typical combination of sensor elements Embodiments of the present invention further include sensors comprising the polycarbonate polymer compositions disclosed herein combined with other sensor elements, such as interference suppression films (e.g., interference suppression films disclosed in U.S. Patent Application Serial No. 12 / 572,087, the contents of which are incorporated herein by reference). One such embodiment of the invention is an interference suppression film comprising a methacrylic acid polymer with a molecular weight between 100 kilodaltons and 1000 kilodaltons, wherein the methacrylic acid polymer is crosslinked by a hydrophilic crosslinking agent such as an organically functional dipodalalkoxysilane. Another embodiment of the invention is an interference suppression film comprising a primary amine polymer with a molecular weight between 4,000 kilodaltons and 500 kilodaltons, wherein the primary amine polymer is crosslinked by a hydrophilic crosslinking agent such as glutaraldehyde. Typically, these interference suppression films are coated with a hydrogen peroxide conversion composition. In illustrative embodiments, the hydrogen peroxide conversion composition comprises an electrode; and the crosslinked interference suppression film is coated on the electrode in a layer with a thickness of 0.1 μm to 1.0 μm.

[0062] In some embodiments of the invention, elements of the sensor device, such as electrodes or orifices, are designed with specific configurations and / or made of specific materials and / or positioned relative to other elements to facilitate sensor functionality. In one such embodiment of the invention, the working electrode, counter electrode, and reference electrode are positioned on a substrate and / or conductive layer in a configuration that facilitates sensor startup and / or maintains hydration of the working electrode, counter electrode, and / or reference electrode when the sensor device is placed in contact with a fluid containing an analyte (e.g., by suppressing electrode shielding, a phenomenon that can suppress hydration and capacitive startup of the sensor circuitry). Generally, such embodiments of the invention facilitate sensor startup and / or initialization.

[0063] Optionally, embodiments of the device include multiple working electrodes and / or counter electrodes and / or reference electrodes (e.g., three working electrodes, one reference electrode, and one counter electrode) to provide, for example, redundant sensing capability. Some embodiments of the invention include a single sensor. Other embodiments of the invention include multiple sensors. In some embodiments of the invention, a pulsed voltage is used to obtain a signal from one or more electrodes of the sensor. Optionally, the multiple working electrodes, counter electrodes, and reference electrodes are configured together as units and distributed in position on the conductive layer in a repeating unit pattern. In some embodiments of the invention, the elongated substrate layer is made of a flexible material that allows the sensor to twist and bend when implanted in vivo; and the electrodes are grouped in a configuration that facilitates intracellular fluid contact with at least one of the working electrodes when twisting and bending occurs during implantation of the sensor device in vivo. In some embodiments, the electrodes are grouped in a configuration that allows the sensor to continue operating if a portion of the sensor having one or more electrodes is removed from the in vivo environment and exposed to an ex vivo environment.

[0064] In some embodiments of the invention that include multiple sensors, elements such as sensor electrodes are organized / arranged within a flexible circuit assembly. In such embodiments of the invention, the architecture of the sensor system can be designed such that the first sensor does not affect signals generated by the second sensor (and vice versa); and that the first and second sensors sense from separate tissue envelopes; therefore, signals from individual sensors do not interact. Meanwhile, in typical embodiments of the invention, the sensors are spaced apart from each other, allowing for easy encapsulation together and / or suitability for implantation via a single insertion action. One such embodiment of the present invention is a device for monitoring analytes in a patient, the device comprising: a substrate element adapted to fix the device to a patient; a first puncture member coupled to and extending from the substrate element; a first electrochemical sensor operatively coupled to the first puncture member and including a first electrochemical sensor electrode for determining at least one physiological characteristic of the patient at a first electrochemical sensor placement site; a second puncture member coupled to and extending from the substrate element; and a second electrochemical sensor operatively coupled to the second puncture member and including a second electrochemical sensor electrode for determining at least one physiological characteristic of the patient at a second electrochemical sensor placement site. In such embodiments of the invention, at least one physiological characteristic monitored by the first or second electrochemical sensor includes the concentration of an analyte naturally present in the patient's body; the first puncture member places the first electrochemical sensor in a first tissue compartment of the patient, and the second puncture member places the second electrochemical sensor in a second tissue compartment of the patient; and the first and second puncture members are disposed on a substrate in a configuration selected to avoid alteration of the sensor signal generated by the second electrochemical sensor due to possible changes in the physiological response caused by the implantation of the first electrochemical sensor.

[0065] Various elements of a sensor device may be positioned at certain locations within the device and / or configured in a specific shape and / or constructed from specific materials to enhance the strength and / or functionality of the sensor. One embodiment of the invention includes an elongated substrate comprising a polyimide or dielectric ceramic material that promotes the strength and durability of the sensor. In some embodiments of the invention, the structural features and / or relative positions of the working electrode and / or counter electrode and / or reference electrode are designed to influence the manufacture, use, and / or functionality of the sensor. Optionally, the sensor is operatively coupled to a series of elements including flexible circuitry (e.g., electrodes, electrical conduits, contact pads, etc.). One embodiment of the invention includes electrodes with one or more rounded edges to suppress delamination of layers disposed on the electrodes (e.g., an analyte sensing layer including glucose oxidase).

[0066] In some embodiments of the invention, the electrodes of the device comprise a platinum composition, and the device further comprises a titanium composition disposed between an elongated substrate layer and a conductive layer. Optionally, in such embodiments, the device further comprises a gold composition disposed between the titanium composition and the conductive layer. Such embodiments of the invention generally exhibit enhanced bonding and / or less corrosion and / or improved biocompatibility profiles between layered materials within the sensor. Related embodiments of the invention include methods for suppressing corrosion of sensor elements and / or methods for improving the biocompatibility of sensor embodiments of the invention (e.g., sensors configured to use such materials).

[0067] In typical embodiments of the invention, the sensor is operatively coupled to other elements (e.g., electronic components), such as elements designed to transmit and / or receive signals, monitors, processors, etc., and devices that can use sensor data to modulate patient physiological functions, such as drug infusion pumps. For example, in some embodiments of the invention, the sensor is operatively coupled to: a sensor input capable of receiving signals from the sensor based on physiological characteristics sensed in mammals; and a processor coupled to the sensor input, wherein the processor is capable of characterizing one or more signals received from the sensor. A variety of sensor configurations as disclosed herein can be used in such systems. Optionally, for example, the sensor includes three working electrodes, a counter electrode, and a reference electrode. In some embodiments, at least one working electrode is coated with an analyte sensing layer including glucose oxidase, and at least one working electrode is not coated with an analyte sensing layer including glucose oxidase.

[0068] Diagrammatic view of typical sensor configuration Figure 1A cross-section of a typical sensor embodiment 100 of the present invention is shown. This sensor embodiment is formed of multiple components, which are generally in the form of layers of various conductive and non-conductive components disposed on each other according to methods recognized in the art and / or specific methods of the present invention disclosed herein. The components of the sensor are generally characterized herein as layers because, for example, they allow for easy characterization. Figure 1 The sensor structure shown is illustrated. However, those skilled in the art will understand that in some embodiments of the invention, the sensor components are combined such that multiple components form one or more heterogeneous layers. In this context, those skilled in the art will understand that the order of the layered components can be varied in various embodiments of the invention.

[0069] Figure 1 The illustrated embodiment includes a substrate layer 102 for supporting the sensor 100. The substrate layer 102 may be made of a material such as a metal and / or ceramic and / or polymer substrate, and may be self-supporting or further supported by another material known in the art. Embodiments of the invention include a conductive layer 104 disposed on and / or combined with the substrate layer 102. Typically, the conductive layer 104 includes one or more electrodes. The operating sensor 100 typically includes multiple electrodes, such as a working electrode, a counter electrode, and a reference electrode. Other embodiments may also include multiple working electrodes and / or counter electrodes and / or reference electrodes and / or one or more electrodes performing multiple functions, such as electrodes that serve as both a reference electrode and a counter electrode.

[0070] As discussed in detail below, many known techniques and materials can be used to produce the base layer 102 and / or conductive layer 104. In some embodiments of the invention, the circuitry of the sensor is defined by etching the conductive layer 104 into a desired conductive path pattern. A typical circuitry for the sensor 100 includes two or more adjacent conductive paths having a region at a proximal end to form a contact pad and a region at a distal end to form a sensor electrode. An electrically insulating cover layer 106, such as a polymer coating, may be disposed on portions of the sensor 100. Acceptable polymer coatings used as insulating protective cover layer 106 may include, but are not limited to, non-toxic biocompatible polymers such as silicone compounds, polyimides, biocompatible solder masks, epoxy acrylate copolymers, etc. In the sensor of the present invention, one or more exposed areas or holes 108 may be formed through the cover layer 106 to open the conductive layer 104 to the external environment and, for example, allow analytes such as glucose to pass through the sensor layer and be sensed by the sensing element. The holes 108 may be formed by a variety of techniques, including laser ablation, band masking, chemical polishing or etching, or photolithography. In some embodiments of the invention, during manufacturing, a second photoresist may be applied to the protective layer 106 to define the area of ​​the protective layer to be removed, thereby forming one or more holes 108. The exposed electrodes and / or contact pads may also undergo secondary processing (e.g., through the holes 108) such as additional electroplating to prepare surfaces and / or reinforce conductive areas.

[0071] exist Figure 1 In the sensor configuration shown, the analyte sensing layer 110 (which is typically a sensor chemistry layer, meaning that the materials in this layer undergo a chemical reaction to generate a signal that can be sensed by the conductive layer) is disposed on one or more exposed electrodes of the conductive layer 104. Figure 2BIn the illustrated sensor configuration, an interference suppression membrane 120 is disposed on one or more exposed electrodes of the conductive layer 104, wherein the analyte sensing layer 110 is then disposed on this interference suppression membrane 120. Typically, the analyte sensing layer 110 is an enzyme layer. Most typically, the analyte sensing layer 110 includes an enzyme capable of generating and / or utilizing oxygen and / or hydrogen peroxide, such as glucose oxidase. Optionally, the enzyme in the analyte sensing layer is combined with a second carrier protein such as human serum albumin, bovine serum albumin, etc. In an illustrative embodiment, an oxidoreductase such as glucose oxidase in the analyte sensing layer 110 reacts with glucose to generate hydrogen peroxide, which subsequently modulates the current at the electrode. Since this modulation of the current depends on the concentration of hydrogen peroxide, and the concentration of hydrogen peroxide is related to the concentration of glucose, the glucose concentration can be determined by monitoring this modulation of the current. In a specific embodiment of the invention, hydrogen peroxide is oxidized at a working electrode serving as the anode (also referred to herein as the anodic working electrode), wherein the resulting current is proportional to the hydrogen peroxide concentration. This modulation of the current caused by the change in hydrogen peroxide concentration can be monitored by any of a variety of sensor detector devices, such as general-purpose sensor ampere-type biosensor detectors or one of the various other similar devices known in the art, such as the glucose monitoring device manufactured by Medtronic MiniMed.

[0072] In embodiments of the invention, the analyte sensing layer 110 may be applied over a portion of or over the entire area of ​​the conductive layer. Typically, the analyte sensing layer 110 is disposed on a working electrode, which may be an anode or a cathode. Optionally, the analyte sensing layer 110 is also disposed on a counter electrode and / or a reference electrode. While the thickness of the analyte sensing layer 110 can be up to about 1000 micrometers (μm), it is generally relatively thin compared to the thicknesses found in sensors previously described in the art, and the thickness is typically, for example, less than 1 micrometer, 0.5 micrometers, 0.25 micrometers, or 0.1 micrometers. Some methods for producing the thin analyte sensing layer 110, as discussed in detail below, include: brushing the layer onto a substrate (e.g., the reactive surface of a platinum black electrode), as well as spin coating, dip coating and drying processes, low-shear spraying processes, inkjet printing processes, screen printing processes, etc.

[0073] Typically, the analyte sensing layer 110 is coated and / or disposed near one or more additional layers. Optionally, one or more additional layers include a protein layer 116 disposed on the analyte sensing layer 110. Typically, the protein layer 116 contains proteins such as human serum albumin, bovine serum albumin, etc. Typically, the protein layer 116 includes human serum albumin. In some embodiments of the invention, an additional layer includes an analyte modulation layer 112 disposed above the analyte sensing layer 110 to modulate the proximity of the analyte to the analyte sensing layer 110. For example, the analyte modulation membrane layer 112 may include a glucose restriction membrane that modulates the amount of glucose in contact with an enzyme (such as glucose oxidase) present in the analyte sensing layer. Such glucose restriction membranes can be made of a variety of materials known to be suitable for such purposes, such as silicone compounds like polydimethylsiloxane, polyurethane, polyurea cellulose acetate, perfluorosulfonic acid (NAFION), polyester sulfonic acid (e.g., Kodak AQ), hydrogels, polymer blends disclosed herein, or any other suitable hydrophilic membranes known to those skilled in the art.

[0074] In some embodiments of the present invention, such as Figure 1 As shown, the adhesion promoter layer 114 is disposed between layers such as the analyte modulation layer 112 and the analyte sensing layer 110 to promote their contact and / or adhesion. In a specific embodiment of the invention, such as Figure 1 As shown, an adhesion promoter layer 114 is disposed between the analyte modulation layer 112 and the protein layer 116 to promote their contact and / or adhesion. The adhesion promoter layer 114 can be made of any of a variety of materials known in the art to promote bonding between such layers. The adhesion promoter layer 114 typically comprises a silane compound. In alternative embodiments, the proteins or similar molecules in the analyte sensing layer 110 can be sufficiently cross-linked or otherwise prepared to allow the analyte modulation membrane layer 112 to be disposed in direct contact with the analyte sensing layer 110 in the absence of the adhesion promoter layer 114.

[0075] Examples of typical components used in manufacturing the sensors disclosed herein are discussed below.

[0076] Typical analyte sensor components used in embodiments of the present invention The following disclosure provides examples of typical elements / components used in sensor embodiments of the present invention. While these elements may be described as discrete units (e.g., layers), those skilled in the art will understand that sensors can be designed to contain some or all of the material properties and / or functions of the elements / components discussed below (e.g., elements acting as a supporting substrate component and / or a conductive component and / or a matrix for both analyte sensing components and further serving as electrodes in the sensor). Those skilled in the art will understand that these thin-film analyte sensors are applicable to many sensor systems, such as those described below.

[0077] Base components The sensors of this invention typically contain a substrate component (see, for example...) Figure 1 (Component 102 in the document). The term "substrate composition" is used herein according to the technical terms recognized in the art and refers to a component in a device that typically provides a support matrix for multiple components stacked on top of each other and including functional sensors. In one form, the substrate composition comprises a thin film of insulating (e.g., electrically insulating and / or waterproof) material. This substrate composition can be made of a variety of materials having desired qualities such as dielectric properties, waterproofness, and hermeticity. Some materials include metals and / or ceramics and / or polymer substrates, etc.

[0078] The substrate component can be self-supporting or can be further supported by another material known in the art. Figure 1 In one embodiment of the sensor configuration shown, substrate composition 102 comprises ceramic. Alternatively, the substrate composition comprises a polymeric material, such as polyamide. In one illustrative embodiment, the ceramic substrate comprises a composition primarily composed of Al2O3 (e.g., 96%). The use of alumina as an insulating substrate composition for use with implantable devices is disclosed in U.S. Patents 4,940,858, 4,678,868, and 6,472,122, which are incorporated herein by reference. The substrate composition of the present invention may further comprise other elements known in the art, such as sealing vias (see, for example, WO 03 / 023388). Depending on the specific sensor design, the substrate composition may be a relatively thick composition (e.g., thicker than 50, 100, 200, 300, 400, 500, or 1000 micrometers). Alternatively, non-conductive ceramics, such as alumina, can be utilized in thin compositions, for example, less than about 30 micrometers.

[0079] Conductive components The electrochemical sensor of the present invention typically comprises a conductive component disposed on a substrate component, including at least one electrode for measuring the analyte or its byproducts (e.g., oxygen and / or hydrogen peroxide) (see example...). Figure 1(Element 104 in the text). The term "conductive component" is used herein according to terminology recognized in the art and refers to a conductive sensor element, such as an electrode, capable of measuring a detectable signal and transmitting said signal to a detection device. An illustrative example of a conductive component is one that, compared to a reference electrode, does not experience changes in the concentration of an analyte (e.g., oxygen, a co-reactant used when the analyte interacts with a composition present in analyte sensing component 110 (e.g., the enzyme glucose oxidase) or a reaction product of such interaction (e.g., hydrogen peroxide)), the conductive component can measure an increase or decrease in current in response to exposure to a stimulus (such as a change in the concentration of the analyte or its byproducts). Illustrative examples of such elements include electrodes capable of generating a variable detectable signal in the presence of molecules (such as hydrogen peroxide or oxygen) at variable concentrations. Typically, one of these electrodes in a conductive component is a working electrode, which may be made of a non-corrosive metal or carbon. Carbon working electrodes may be vitreous or graphitic and may be made of a solid or a paste. The metal working electrode can be made of platinum group metals, including palladium or gold, or of a non-corrosive conductive metal oxide, such as ruthenium dioxide. Alternatively, the electrode can comprise a silver / silver chloride electrode composition. The working electrode can be, for example, a wire or a thin conductive film applied to a substrate by coating or printing. Typically, only a portion of the surface of the metal or carbon conductor is electrolytically in contact with the solution containing the analyte. This portion is referred to as the working surface of the electrode. The remaining surface of the electrode is typically isolated from the solution by an electrically insulating covering component 106. Examples of useful materials for producing this protective covering component 106 include polymers such as polyimides, polytetrafluoroethylene, polyhexafluoropropylene, and silicones such as polysiloxanes.

[0080] In addition to the working electrode, the analyte sensor of the present invention typically includes a reference electrode or a combination of a reference electrode and a counter electrode (also referred to as a quasi-reference electrode or counter / reference electrode). If the sensor does not have a counter / reference electrode, it may include a separate counter electrode, which may be made of the same or different material as the working electrode. A typical sensor of the present invention has one or more working electrodes and one or more counter, reference, and / or counter / reference electrodes. One embodiment of the sensor of the present invention has two, three, four, or more working electrodes. These working electrodes in the sensor may be integrally connected or may remain separate.

[0081] Typically, for in vivo use, embodiments of the present invention are subcutaneously implanted into the skin of a mammal to directly contact the mammal's bodily fluids, such as blood. Alternatively, the sensor can be implanted in other areas of the mammalian body, such as the space within the peritoneum. When multiple working electrodes are used, the multiple working electrodes can be implanted together or at different locations within the body. Counter electrodes, reference electrodes, and / or counter / reference electrodes can also be implanted near one or more working electrodes or at other locations within the mammalian body. Embodiments of the present invention include sensors comprising electrodes made of nanostructured materials. As used herein, "nanostructured material" is an object fabricated to have a size of at least one less than 100 nm. Examples include, but are not limited to, single-walled nanotubes, double-walled nanotubes, multi-walled nanotubes, nanotube bundles, fullerenes, cocoons, nanowires, nanofibers, onions, etc.

[0082] Interference inhibition components The electrochemical sensor of the present invention optionally includes an interference-suppressing component disposed between the electrode surface and the environment to be measured. Specifically, some sensor embodiments rely on the oxidation and / or reduction of hydrogen peroxide produced by an enzymatic reaction on the surface of the working electrode under a constant applied potential. Because amperometric detection based on the direct oxidation of hydrogen peroxide requires a relatively high oxidation potential, sensors employing this detection scheme can experience interference from oxidizable substances present in biological fluids, such as ascorbic acid, uric acid, and acetamide phenol. In this context, the term "interference-suppressing component" is used herein according to recognized technical terms in the art and refers to a coating or film in the sensor that serves to suppress stray signals generated by such oxidizable species that interfere with the detection of signals generated by the analyte to be sensed. Some interference-suppressing components function through size exclusion (e.g., by excluding interfering species of a specific size). Examples of interference-suppressing components include one or more compound layers or coatings, such as hydrophilic crosslinked pHEMA and lysine polymers, and cellulose acetate (including cellulose acetate incorporating reagents such as poly(ethylene glycol), polyethersulfone, polytetrafluoroethylene, perfluorinated ionomer NAFION, polyphenylene diamine, epoxy resins, etc.). Illustrative discussions of such interference-suppressing components can be found, for example, in Ward et al., Biosensors and Bioelectronics 17 (2002) 181-189 and Choi et al., Analytical Chimica Acta 461 (2002) 251-260, which are incorporated herein by reference. Other interference-suppressing components include, for example, those observed to restrict compound movement based on molecular weight range, such as cellulose acetate disclosed, for example, in U.S. Patent No. 5,755,939, the contents of which are incorporated herein by reference. This document discloses, for example, in U.S. Patent Application Serial No. 12 / 572,087, a composition having a series of undesirable material properties that could make other compositions ideally used as interference suppression membranes in certain amperometric glucose sensors, and methods of manufacturing and using thereof.

[0083] Analyte Sensing Components The electrochemical sensor of the present invention comprises an analyte sensing component disposed on an electrode of the sensor (see, for example...). Figure 1(Element 110 in the text). The term "analyte sensing component" is used herein according to recognized technical terms in the art and refers to a component comprising a material capable of recognizing or reacting with an analyte to be detected by an analyte sensor device. Typically, this material in the analyte sensing component typically generates a detectable signal via an electrode of the conductive component upon interaction with the analyte to be sensed. In this respect, the electrodes of the analyte sensing component and the conductive component operate in combination to generate an electrical signal that can be read by a device associated with the analyte sensor. Typically, the analyte sensing component includes an oxidoreductase (e.g., glucose oxidase) capable of reacting with molecules and / or causing a concentration change of said molecules, the concentration change of said molecules can be measured by measuring a change in current at an electrode of the conductive component (e.g., oxygen and / or hydrogen peroxide). Enzymes capable of generating molecules (e.g., hydrogen peroxide) can be disposed on electrodes according to many methods known in the art. The analyte sensing component can coat all or part of the respective electrodes of the sensor. In this context, the analyte sensing component can coat the electrodes to an equivalent extent. Alternatively, the analyte sensing component can be coated to varying degrees on different electrodes, wherein, for example, the coated surface of the working electrode is larger than the coated surface of the counter electrode and / or reference electrode.

[0084] A typical sensor embodiment of this element of the present invention utilizes an enzyme (e.g., glucose oxidase) that has been combined in a fixed proportion with a second protein (e.g., albumin) and then applied to the surface of the electrode to form a thin enzyme component (e.g., an enzyme typically optimized for the stability properties of glucose oxidase). In a typical embodiment, the analyte sensing component comprises a mixture of GOx and HSA. In a typical embodiment of the analyte sensing component containing GOx, GOx reacts with glucose present in the sensing environment (e.g., the body of a mammal) and according to... Figure 1 The reaction shown produces hydrogen peroxide, which is thus produced and detected at the anode at the working electrode in the conductive component.

[0085] As described above, enzymes and second proteins (e.g., albumin) are typically treated to form a cross-linked matrix, for example, by adding a cross-linking agent to a protein mixture. As is known in the art, cross-linking conditions can be manipulated to modulate factors such as the retention of the enzyme's biological activity, its mechanical and / or operational stability. Illustrative cross-linking procedures are described in U.S. Patent Application Serial No. 10 / 335,506 and PCT Publication WO 03 / 035891, which are incorporated herein by reference. For example, an amine cross-linking agent (such as, but not limited to, glutaraldehyde) can be added to the protein mixture.

[0086] protein components The electrochemical sensor of the present invention optionally includes a protein component disposed between the analyte sensing component and the analyte modulation component (see, for example...). Figure 1 (Element 116 in the text). The term "protein component" is used herein according to recognized technical terms in the art and refers to a component containing carrier proteins, etc., selected to be compatible with the analyte sensing component and / or analyte modulation component. In a typical embodiment, the protein component includes albumin, such as human serum albumin. The HSA concentration can vary between about 0.5% and 30% (w / v). Typically, the HSA concentration is about 1-10% w / v, and most typically about 5% w / v. In alternative embodiments of the invention, collagen or BSA or other structural proteins used in these contexts may be used in place of or to supplement the HSA. According to recognized practices in the art, this component is typically cross-linked onto the analyte sensing component.

[0087] Adhesion promoting ingredients The electrochemical sensor of the present invention may contain one or more adhesion-promoting (AP) components (see, for example) Figure 1 (Element 114 in the text). The term "adhesion-promoting component" is used herein according to recognized technical terms in the art and refers to a component comprising a material selected to promote adhesion between adjacent components in a sensor. Adhesion-promoting components are typically disposed between an analyte-sensing component and an analyte-modifying component. Adhesion-promoting components are typically disposed between an optional protein component and an analyte-modifying component. Adhesion-promoting components can be made from any of a variety of materials known in the art for promoting adhesion between such components and can be applied by any of a variety of methods known in the art. Adhesion-promoting components typically include silane compounds, such as γ-aminopropyltrimethoxysilane.

[0088] The use of silane coupling agents, particularly silane coupling agents of the formula R'Si(OR)3, wherein R' is typically an aliphatic group having a terminal amine and R is a lower alkyl group, to promote adhesion is known in the art (see, for example, U.S. Patent No. 5,212,050, which is incorporated herein by reference). For example, silanes such as γ-aminopropyltriethoxysilane and glutaraldehyde are used in stepwise methods to induce bovine serum albumin (BSA) and glucose oxidase (GO). X Chemically modified electrodes that adhere to and co-crosslink with the electrode surface are known in the art (see, for example, Yao, T. Journal of Analytical Chemistry, 1983, 148, 27-33).

[0089] In some embodiments of the invention, the adhesion-promoting component further includes one or more compounds that may also be present in adjacent components, such as polydimethylsiloxane (PDMS) compounds, for limiting the diffusion of analytes, such as glucose, through the analyte modulation component. In illustrative embodiments, the formulation comprises 0.5-20% PDMS, typically 5-15% PDMS, and most typically 10% PDMS. In some embodiments of the invention, the adhesion-promoting component is cross-linked within a layered sensor system and correspondingly comprises an agent selected for its ability to partially cross-link portions present in proximal components, such as the analyte modulation component. In illustrative embodiments of the invention, the adhesion-promoting component comprises an agent selected for its ability to cross-link amine or carboxyl portions of proteins present in proximal components, such as the analyte sensing component and / or protein components, and / or to partially cross-link siloxanes present in compounds disposed in proximal layers, such as the analyte modulation layer.

[0090] Analyte modulation components The electrochemical sensor of the present invention includes an analyte modulation component disposed on the sensor (see, for example...). Figure 1 (Element 112 in the document). The term "analyte modulation component" is used herein according to the technically accepted terminology in the art and refers to a component that typically forms a membrane on a sensor, said membrane operating to modulate the diffusion of one or more analytes (such as glucose) through said component. In some embodiments of the invention, the analyte modulation component is an analyte-restricting membrane (e.g., a glucose-restricting membrane) that operates to block or restrict the diffusion of one or more analytes (such as glucose) through said component. In other embodiments of the invention, the analyte modulation component operates to facilitate the diffusion of one or more analytes through said component. Optionally, such analyte modulation components may be formed to block or restrict the diffusion of one type of molecule (e.g., glucose) through said component, while allowing or even facilitating the diffusion of other types of molecules (e.g., O2) through said component. Typically, the analyte modulation component comprises a polycarbonate polymer composition as disclosed herein.

[0091] Regarding glucose sensors, in known enzyme electrodes, glucose and oxygen from the blood, along with some interfering substances (such as ascorbic acid and uric acid), diffuse through the primary membrane of the sensor. When glucose, oxygen, and interfering substances reach the analyte detection component, an enzyme (such as glucose oxidase) catalyzes the conversion of glucose into hydrogen peroxide and gluconolactone. The hydrogen peroxide can diffuse back through the analyte modulation component, or it can diffuse to the electrode, where it can react to form oxygen and protons to generate a current proportional to the glucose concentration. The sensor membrane assembly has multiple functions, including selectively allowing glucose to pass through. In this case, the illustrative analyte modulation component is a semi-permeable membrane that allows water, oxygen, and at least one selective analyte to pass through and has water-absorbing capacity; the membrane has a water-soluble hydrophilic polymer.

[0092] Various illustrative analyte modulation compositions are known in the art and described, for example, in U.S. Patent Nos. 6,319,540, 5,882,494, 5,786,439, 5,777,060, 5,771,868, and 5,391,250, the disclosure of each of which is incorporated herein by reference. The hydrogels described therein are particularly suitable for use with various implantable devices, for which it is advantageous to provide a surrounding water component. In a typical embodiment of the invention, the analyte modulation composition comprises the polycarbonate polymer composition disclosed herein.

[0093] Covering ingredients The electrochemical sensor of the present invention comprises one or more covering components, typically electrically insulating and protective components (see, for example...). Figure 1 (element 106 in the text). Typically, such covering components may be in the form of a coating, sheath, or tube and are disposed on at least a portion of the analyte modulation component. Acceptable polymer coatings used as insulating protective covering components may comprise, but are not limited to, non-toxic biocompatible polymers such as silicone compounds, polyimides, biocompatible welding masks, epoxy acrylate copolymers, etc. Further, these coatings may be photoimageable to facilitate photolithographic formation of apertures through the conductive component. Typical covering components include spinning on silicone. As is known in the art, this component may be a commercially available RTV (room temperature vulcanizing) silicone composition. In this context, a typical chemical is polydimethylsiloxane (acetoxy-based).

[0094] Illustrative Examples of Analytical Sensor Devices and Associated Characteristics The analyte sensor devices disclosed herein have many embodiments. A general embodiment of the invention is an analyte sensor device for implantation in mammals. While analyte sensors are generally designed to be implanted in mammals, the sensors are not limited to any particular environment but can be used in a variety of situations, such as for analyzing most liquid samples, including biological fluids such as whole blood, lymph, plasma, serum, saliva, urine, feces, sweat, mucus, tears, cerebrospinal fluid, nasal secretions, cervical or vaginal secretions, semen, pleural fluid, amniotic fluid, peritoneal fluid, middle ear effusion, synovial fluid, gastric juice, etc. Additionally, solid or dry samples can be dissolved in a suitable solvent to provide a liquid mixture suitable for analysis.

[0095] As described above, the sensor embodiments disclosed herein can be used to sense analytes of interest in one or more physiological environments. For example, in some embodiments, as is typically the case with subcutaneous sensors, the sensor may be in direct contact with interstitial fluid. The sensor of the present invention can also be part of a skin surface system in which interstitial glucose is extracted through the skin and brought into contact with the sensor (see, for example, U.S. Patent Nos. 6,155,992 and 6,706,159, which are incorporated herein by reference). In other embodiments, as is typically the case with intravenous sensors, the sensor may be in contact with blood. The sensor embodiments of the present invention further include sensors suitable for a variety of situations. For example, in some embodiments, the sensor may be designed for use in mobile situations, such as those employed by mobile users. Alternatively, the sensor may be designed for use in stationary situations, such as those suitable for clinical settings. Such sensor embodiments include, for example, sensors for monitoring one or more analytes present in one or more physiological environments in hospitalized patients.

[0096] The sensor of the present invention can also be incorporated into a wide variety of medical systems known in the art. The sensor of the present invention can be used, for example, in closed-loop infusion systems designed to control the rate at which a drug is infused into a user's body. Such closed-loop infusion systems may include a sensor and associated instruments that generate inputs to a controller, which in turn operates the delivery system (e.g., a system for calculating the dose to be delivered by a drug infusion pump). In such cases, the instruments associated with the sensor may also transmit commands to the delivery system and can be used to remotely control the delivery system. Typically, the sensor is a subcutaneous sensor that contacts interstitial fluid to monitor glucose concentration in the user's body, and the fluid infused into the user's body by the delivery system contains insulin. Illustrative systems are disclosed, for example, in U.S. Patent Nos. 6,558,351 and 6,551,276; PCT Applications Nos. 99 / 21703 and 99 / 22993; and WO 2004 / 008956 and WO 2004 / 009161, all of which are incorporated herein by reference.

[0097] Some embodiments of the present invention measure peroxides and have advantageous characteristics suitable for implantation in various sites in mammals, including subcutaneous and intravenous implantation areas, as well as implantation in various non-vascular regions. Due to the oxygen noise problems that can occur in oxygen sensors implanted in non-vascular regions, peroxide sensor designs implanted in non-vascular regions allow for advantages over certain oxygen measurement sensor device designs. For example, in such implantable oxygen sensor device designs, oxygen noise at the reference sensor can impair the signal-to-noise ratio, thereby interfering with the implantable oxygen sensor device's ability to obtain stable glucose readings in such environments. Therefore, the peroxide sensor of the present invention overcomes the difficulties observed with such oxygen sensors in non-vascular regions.

[0098] Some embodiments of the peroxide sensor of the present invention further comprise advantageous long-term or “permanent” sensors suitable for implantation in mammals for a duration greater than 30 days. Specifically, as is known in the art (see, for example, ISO 10993, “Biological Evaluation of Medical Devices”), medical devices with sensors as described herein can be categorized into three groups based on implantation duration: (1) “limited” (< 24 hours), (2) “extended” (24 hours–30 days), and (3) “permanent” (> 30 days). In some embodiments of the present invention, the design of the peroxide sensor of the present invention allows for “permanent” implantation, i.e., > 30 days, according to this classification. In relevant embodiments of the present invention, the highly stable design of the peroxide sensor of the present invention allows the implantable sensor to continue operating for 2, 3, 4, 5, 6 months, or 12 months or more in this respect.

[0099] Arrangement of analyte sensor devices and components As described above, the invention disclosed herein includes numerous embodiments comprising a sensor having a series of elements comprising a polycarbonate polymer membrane. Such embodiments of the invention allow those skilled in the art to produce various arrangements of the analyte sensor devices disclosed herein. As described above, illustrative general embodiments of the sensors disclosed herein comprise a substrate layer, a capping layer, and at least one layer having a sensor element, such as an electrode, disposed between the substrate layer and the capping layer. Typically, exposed portions of one or more sensor elements (e.g., working electrode, counter electrode, reference electrode, etc.) are coated with a very thin layer of material having a suitable electrode chemical substance. For example, enzymes such as lactate oxidase, glucose oxidase, glucose dehydrogenase, or hexokinase may be disposed on exposed portions of the sensor element located within openings or pores defined in the capping layer. Figure 1 A cross-section of a typical sensor structure 100 of the present invention is shown. According to the method for producing the sensor structure 100 of the present invention, the sensor is formed from multiple layers of various conductive and non-conductive components disposed on each other.

[0100] As described above, in the sensor of the present invention, the layers of the sensor (e.g., the analyte sensing layer) may have one or more bioactive materials and / or inert materials incorporated therein. The term "incorporated" as used herein refers to any state or condition describing the incorporated material being held on the outer surface of the layer or within the solid phase or supporting matrix of the layer. Thus, the "incorporated" material may, for example, be fixed, physically encapsulated, or covalently adhered to functional groups of one or more matrix layers. Furthermore, any additional steps or agents that promote the "incorporation" of said material may be employed if they are not detrimental to the present invention but are consistent with its objectives. Of course, this definition applies to any embodiment of the present invention in which bioactive molecules (e.g., enzymes such as glucose oxidase) are "incorporated." For example, some layers of the sensor disclosed herein contain protein material, such as albumin, as used as a crosslinkable matrix. As used herein, protein material means encompassing substances generally derived from proteins, whether the actual substance is a natural protein, an inactivated protein, a denatured protein, a hydrolyzed substance, or a derivative thereof. Examples of suitable protein materials include, but are not limited to, enzymes such as glucose oxidase and lactate oxidase, albumins (e.g., human serum albumin, bovine serum albumin, etc.), casein, gamma globulin, collagen and collagen derivatives (e.g., fish gelatin, fish glue, animal gelatin and animal glue).

[0101] Figure 1 An illustrative embodiment of the invention is shown below. This embodiment includes an electrically insulating substrate layer 102 for supporting a sensor 100. The electrically insulating substrate layer 102 may be made of a material such as a ceramic substrate, and may be self-supporting or may be further supported by another material known in the art. In an alternative embodiment, the electrically insulating layer 102 comprises a polyimide substrate, such as a polyimide tape, dispensed from a reel. Providing layer 102 in this form facilitates clean, high-density mass production. Furthermore, in some production processes using such polyimide tapes, the sensor 100 may be produced on both sides of the tape.

[0102] A typical embodiment of the present invention includes an analyte sensing layer disposed on a substrate 102. In such... Figure 1In the illustrative embodiment shown, the analyte sensing layer includes a conductive layer 104 disposed on an insulating substrate layer 102. Typically, the conductive layer 104 includes one or more electrodes. As described below, the conductive layer 104 can be applied using many known techniques and materials; however, the circuitry of the sensor 100 is typically defined by etching the disposed conductive layer 104 into a desired conductive path pattern. A typical circuitry for the sensor 100 includes two or more adjacent conductive paths having a region at a proximal end to form a contact pad and a region at a distal end to form a sensor electrode. An electrically insulating protective overlay 106, such as a polymer coating, can typically be disposed on portions of the conductive layer 104. Acceptable polymer coatings used as the insulating protective layer 106 can include, but are not limited to, non-toxic, biocompatible polymers such as polyimides, biocompatible welding masks, epoxy acrylate copolymers, etc. Furthermore, these coatings can be photoimageable to facilitate photolithographic formation of apertures 108 through the conductive layer 104. In some embodiments of the invention, the analyte sensing layer is disposed on a porous metal and / or ceramic and / or polymer matrix, wherein this combination of elements serves as an electrode in a sensor.

[0103] In the sensor of the present invention, one or more exposed areas or holes 108 may be formed through the protective layer 106 to the conductive layer 104 to define the contact pads and electrodes of the sensor 100. The holes 108 may be formed by a variety of techniques, including laser ablation, chemical polishing, or etching, in addition to photolithography. A second photoresist may also be applied to the capping layer 106 to define the area to be removed from the protective layer to form the holes 108. The operating sensor 100 typically includes multiple electrodes, such as a working electrode and a counter electrode, that are electrically insulated from each other, however, are typically positioned very close to each other. Other embodiments may also include a reference electrode. Still other embodiments may utilize a separate reference element not formed on the sensor. The exposed electrodes and / or contact pads may also undergo secondary processing through the holes 108, such as additional electroplating, to prepare surfaces and / or reinforce conductive areas.

[0104] An analyte sensing layer 110 is typically disposed through a hole 108 on one or more exposed electrodes of the conductive layer 104. The analyte sensing layer 110 is typically a sensor chemistry layer and most commonly an enzyme layer. The analyte sensing layer 110 typically includes glucose oxidase or lactate oxidase. In such embodiments, the analyte sensing layer 110 reacts with glucose to generate hydrogen peroxide that modulates a current to the electrode, which can be monitored to measure the amount of glucose present. The sensor chemistry layer 110 can be applied over a portion of the conductive layer or over the entire area of ​​the conductive layer. The sensor chemistry layer 110 is typically disposed on portions of the working electrode and counter electrode that include the conductive layer. Some methods for producing a thin sensor chemistry layer 110 include spin coating, dip and dry processes, low-shear spraying, inkjet printing, screen printing, etc. Most commonly, a spin coating process is used to apply a thin sensor chemistry layer 110.

[0105] The analyte sensing layer 110 is typically coated with one or more coatings. In some embodiments of the invention, one such coating comprises a membrane that modulates the amount of analyte that can contact the enzyme in the analyte sensing layer. For example, the coating may comprise an analyte modulation membrane layer, such as a glucose restriction membrane, which modulates the amount of glucose that contacts the glucose oxidase layer at the electrode. Such glucose restriction membranes can be made from a wide variety of materials known to be suitable for this purpose, such as silicone, polyurethane, polyurea cellulose acetate, perfluorosulfonic acid, polyester sulfonic acid (Kodak AQ), hydrogels, or any other membrane known to those skilled in the art. In some embodiments of the invention, the analyte modulation layer comprises a linear polyurethane / polyurea polymer polycarbonate having a branched acrylate hydrophilic comb copolymer having a central chain and a plurality of side chains coupled to the central chain, wherein at least one side chain comprises a silicone moiety.

[0106] In some embodiments of the present invention, the coating is a glucose restriction film layer 112 disposed above the sensor chemical layer 110 to regulate the contact between glucose and the sensor chemical layer 110. In some embodiments of the present invention, such as... Figure 1 As shown, an adhesion promoter layer 114 is disposed between the membrane layer 112 and the sensor chemical layer 110 to promote their contact and / or adhesion. The adhesion promoter layer 114 can be made of any of a variety of materials known in the art to promote bonding between such layers. The adhesion promoter layer 114 typically comprises a silane compound. In alternative embodiments, proteins or similar molecules in the sensor chemical layer 110 may be sufficiently cross-linked or otherwise prepared to allow the membrane layer 112 to be disposed in direct contact with the sensor chemical layer 110 in the absence of the adhesion promoter layer 114.

[0107] As described above, embodiments of the present invention may include one or more functional coatings. As used herein, the term "functional coating" means a layer that coats at least a portion of at least one surface of a sensor, more generally substantially all of the surface of a sensor, and is capable of interacting with one or more analytes, such as compounds, cells, and fragments thereof, in the environment in which the sensor is disposed. Non-limiting examples of functional coatings include sensor chemical layers (e.g., enzyme layers), analyte-restricting layers, biocompatibility layers; layers that increase the slippage of the sensor; layers that promote cell-to-cell connectivity with the sensor; layers that reduce cell-to-cell connectivity with the sensor; and so on. Typically, the analyte-modulating layer operates to prevent or restrict the diffusion of one or more analytes, such as glucose, through the layer. Optionally, such layers may be formed to prevent or restrict the diffusion of one type of molecule (e.g., glucose) through the layer while allowing or even promoting the diffusion of other types of molecules (e.g., O2) through the layer. Illustrative functional coatings are hydrogels, such as those disclosed in U.S. Patent Nos. 5,786,439 and 5,391,250, the disclosures of which are incorporated herein by reference. The hydrogel described therein is particularly suitable for use with a variety of implantable devices, for which it is advantageous to provide a surrounding water layer.

[0108] The sensor embodiments disclosed herein may include a layer having a UV-absorbing polymer. According to one aspect of the invention, a sensor comprising at least one functional coating comprising a UV-absorbing polymer is provided. In some embodiments, the UV-absorbing polymer is a polyurethane, polyurea, or a polyurethane / polyurea copolymer. More generally, the selected UV-absorbing polymer is formed from a reaction mixture comprising a diisocyanate, at least one diol, a diamine, or a mixture thereof, and a multifunctional UV-absorbing monomer.

[0109] UV-absorbing polymers are advantageously used in various sensor manufacturing methods, such as those described below: U.S. Patent No. 5,390,671 to Lord et al. entitled "Transcutaneous Sensor Insertion Set"; U.S. Patent No. 5,165,407 to Wilson et al. entitled "Implantable Glucose Sensor"; and U.S. Patent No. 4,890,620 to Gough et al. entitled "Two-Dimensional Diffusion Glucose Substrate Sensing Electrode," all of which are incorporated herein by reference in their entirety. However, any sensor manufacturing method that includes the step of forming a UV-absorbing polymer layer above or below the sensor element is considered to be within the scope of this invention. Specifically, the methods of this invention are not limited to thin-film manufacturing methods and can work in conjunction with other sensor manufacturing methods utilizing UV laser cutting. Examples can work with thick-film, planar, or cylindrical sensors, as well as other sensor shapes requiring laser cutting.

[0110] As disclosed herein, the sensors of the present invention are specifically designed for use as subcutaneous or transdermal glucose sensors to monitor blood glucose levels in diabetic patients. Typically, each sensor comprises multiple sensor elements, such as conductive elements, like elongated thin-film conductors, formed between a bottom insulating film substrate layer and an overlying insulating film cover layer.

[0111] Multiple different sensor elements can be included in a single sensor if desired. For example, both conductive and reactive sensor elements can be combined in one sensor, optionally with each sensor element disposed on a different portion of the substrate layer. One or more control elements may also be provided. In such embodiments, the sensor may define multiple openings or holes in its cover layer. One or more openings may also be defined directly over a portion of the substrate layer in the cover layer to provide information about the interaction between the substrate layer and one or more analytes in the environment in which the sensor is disposed. The substrate layer and cover layer may comprise a variety of materials, typically polymers. In more specific embodiments, the substrate layer and cover layer comprise insulating materials such as polyimide. Openings are typically formed in the cover layer to expose a distal electrode and a proximal contact pad. For example, in glucose monitoring applications, the sensor may be placed percutaneously such that the distal electrode contacts the patient's blood or extracellular fluid, and the contact pad is disposed externally for easy connection to a monitoring device.

[0112] Analyte sensor equipment configuration In clinical settings, electrochemical sensors can be used to determine accurate and relatively rapid results from blood samples, such as glucose and / or lactate levels. Conventional sensors are manufactured in large quantities, including many servicable components, or in small, planar sensors, which may be more convenient in many cases. As used herein, the term "planar" refers to a well-known procedure, for example, using well-known thick-film or thin-film techniques to fabricate a substantially planar structure comprising a relatively thin layer of material. See, for example, U.S. Patent No. 4,571,292 to Liu et al. and U.S. Patent No. 4,536,274 to Papadakis et al., both of which are incorporated herein by reference. As described below, embodiments of the invention disclosed herein have a wider range of geometric configurations (e.g., planar) than existing sensors in the prior art. Additionally, some embodiments of the invention include one or more sensors disclosed herein coupled to another device, such as a drug infusion pump.

[0113] Figure 2 provides a schematic view of typical analyte sensor configurations of the present invention. Some sensor configurations are presented in a relatively flat "ribbon" shape, which can be manufactured using analyte sensor devices. This "ribbon" shape configuration demonstrates the advantages of the sensors disclosed herein, which arise from the spin coating of sensing enzymes such as glucose oxidase, a manufacturing step that produces an extremely thin enzyme coating that allows for the design and production of highly flexible sensor geometries. Such thin enzyme-coated sensors offer additional advantages, such as allowing for a smaller sensor area while maintaining sensor sensitivity, a highly desirable feature for implantable devices (e.g., smaller devices are easier to implant). Therefore, sensor embodiments of the present invention utilizing very thin analyte sensing layers that can be formed by processes such as spin coating can have a wider range of geometric configurations (e.g., planar) compared to sensors utilizing enzyme layers formed by processes such as electrodeposition.

[0114] Some sensor configurations include multiple conductive elements, such as multiple working electrodes, counter electrodes, and reference electrodes. Advantages of such configurations include increased surface area, resulting in higher sensor sensitivity. For example, one sensor configuration introduces a third working sensor. A significant advantage of this configuration is the averaging of signals from the three sensors, thereby improving sensor accuracy. Other advantages include the ability to measure multiple analytes. Specifically, analyte sensor configurations in this arrangement, including electrodes (e.g., multiple working electrodes, counter electrodes, and reference electrodes), can be combined into multiple analyte sensors. Measurement of multiple analytes, such as oxygen, hydrogen peroxide, glucose, lactate, potassium, calcium, and any other physiologically relevant substances / analytes, offers numerous advantages, such as the ability of such sensors to provide linear responses and ease of calibration and / or recalibration.

[0115] An exemplary multi-sensor device includes a single device having: a first sensor, which is cathode-polarized and designed to measure changes in oxygen concentration at a working electrode (cathode) due to the interaction of glucose with glucose oxidase; and a second sensor, which is anodic-polarized and designed to measure changes in hydrogen peroxide concentration at a working electrode (anode) due to the interaction of glucose from the external environment with glucose oxidase. As is known in the art, in such designs, when oxygen contacts the sensor, the first oxygen sensor typically experiences a decrease in current at the working electrode, while when… Figure 1 When the generated hydrogen peroxide contact sensor is used, the second hydrogen peroxide sensor typically experiences an increase in current at the working electrode. Additionally, as is known in the art, the observation of current changes occurring at the working electrode compared to the reference electrode in a corresponding sensor system is correlated with changes in the concentrations of oxygen and hydrogen peroxide molecules, which may then be correlated with glucose concentrations in the external environment (e.g., the body of a mammal).

[0116] The analyte sensor of the present invention can be coupled to other medical devices such as drug infusion pumps. In an illustrative variant of this invention, the replaceable analyte sensor of the present invention can be coupled to other medical devices such as drug infusion pumps, for example, via a port coupled to the medical device (e.g., a subcutaneous port with a locking electrical connection).

[0117] Illustrative methods and materials for manufacturing the analytical analyte sensor device of the present invention Numerous articles, U.S. patents, and patent applications describe prior art using common methods and materials disclosed herein, and further describe various elements (and methods of their manufacture) that can be used in the sensor designs disclosed herein. These include, for example, U.S. Patent Nos. 6,413,393; 6,368,274; 5,786,439; 5,777,060; 5,391,250; 5,390,671; 5,165,407, 4,890,620, 5,390,671, 5,390,691, 5,391,250, 5,482,473, 5,299,571, and 5,568,806; U.S. Patent Application 20020090738; and PCT International Publications Nos. WO 01 / 58348, WO 03 / 034902, WO 03 / 035117, and WO The contents of each of the following documents are incorporated herein by reference: WO 03 / 035891, WO 03 / 023388, WO 03 / 022128, WO 03 / 022352, WO 03 / 023708, WO 03 / 036255, WO 03 / 036310 and WO 03 / 074107.

[0118] Typical sensors used to monitor glucose concentrations in diabetic patients are further described in Shichiri et al., “In Vivo Characteristics of Needle-Type Glucose Sensor-Measurements of Subcutaneous Glucose Concentrations in Human Volunteers,” Horm. Metab. Res., Supplement 20:17-20 (1988); Bruckel et al., “In Vivo Measurement of Subcutaneous Glucose Concentrations with an Enzymatic Glucose Sensor and a Wick Method,” Klin. Wochenschr., 67:491-495 (1989); and Pickup et al., “In Vivo Molecular Sensing in Diabetes Mellitus: An Implantable Glucose Sensor Utilizing Direct Electron Transfer.” "Implantable Glucose Sensor with Direct Electron Transfer", Diabetologia 32:213-217 (1989). Other sensors are described, for example, Advances in Implantable Devices by Reach et al., A. Turner (ed.), JAI Publishing, London, Chapter 1, (1993), which are incorporated herein by reference.

[0119] A typical embodiment of the invention disclosed herein is a method of manufacturing a sensor device for implantation in mammals, the method comprising the steps of: providing a substrate layer; forming a conductive layer on the substrate layer, wherein the conductive layer comprises electrodes (and typically a working electrode, a reference electrode, and a counter electrode); forming an analyte sensing layer on the conductive layer, wherein the analyte sensing layer comprises a composition that can alter the current at the electrodes in the conductive layer in the presence of an analyte; optionally forming a protein layer on the analyte sensing layer; forming an adhesion-promoting layer on the analyte sensing layer or the optional protein layer; forming an analyte modulation layer disposed on the adhesion-promoting layer, wherein the analyte modulation layer comprises a composition that modulates the diffusion of the analyte through it; and forming a capping layer disposed on at least a portion of the analyte modulation layer, wherein the capping layer further comprises a pore over at least a portion of the analyte modulation layer. In some embodiments of the invention, the analyte modulation layer comprises a linear polyurethane / polyurea polymer polycarbonate having a branched acrylate copolymer having a central chain and a plurality of side chains coupled to the central chain. In some embodiments of these methods, the analyte sensor device is formed in a planar geometry configuration.

[0120] As disclosed herein, the various layers of the sensor can be manufactured to exhibit a variety of different properties, which can be manipulated according to the specific design of the sensor. For example, the adhesion-enhancing layer contains compounds selected for their ability to stabilize the overall sensor structure, typically silane compositions. In some embodiments of the invention, the analyte sensing layer is formed by a spin-coating process and its thickness is selected from the group consisting of: heights less than 1 micrometer, 0.5 micrometer, 0.25 micrometer, and 0.1 micrometer.

[0121] Methods for manufacturing sensors typically include the step of forming a protein layer on an analyte sensing layer, wherein the proteins in the protein layer are albumins selected from the group consisting of bovine serum albumin and human serum albumin. Methods for manufacturing sensors also typically include the step of forming an analyte sensing layer comprising an enzyme composition selected from the group consisting of glucose oxidase, glucose dehydrogenase, lactate oxidase, hexokinase, and lactate dehydrogenase. In such methods, the analyte sensing layer typically comprises a carrier protein composition having a substantially fixed ratio to the enzyme, and the enzyme and carrier protein are distributed in a substantially uniform manner throughout the analyte sensing layer.

[0122] The electrodes of the present invention can be formed from a wide variety of materials known in the art. For example, the electrodes can be made of noble late transition metals. Metals such as gold, platinum, silver, rhodium, iridium, ruthenium, palladium, or osmium are suitable for various embodiments of the invention. In some sensor embodiments, other compositions such as carbon or mercury are also useful. Among these metals, silver, gold, or platinum are commonly used as reference electrode metals. Subsequently chlorinated silver electrodes are commonly used as reference electrodes. These metals can be deposited by any means known in the art, including the plasma deposition methods cited above, or by an electroless method, which may involve depositing the metal onto a previously metallized region while the substrate is immersed in a solution containing a metal salt and a reducing agent. The electroless method continues as the reducing agent contributes electrons to the conductive (metallized) surface, accompanied by the reduction of the metal salt at the conductive surface. The result is the formation of an adsorbed metal layer. (For further discussion on electroless methods, see: Wise, EM, *Palladium: Recovery, Properties, and Uses*, Academic Press, New York, NY, (1988); Wong, K. et al., *Plating and Surface Finishing*, 1988, 75, 70-76; Matsuoka, M. et al., ibid., 1988, 75, 102-106; and Pearlstein, F., “Electroless Plating,” *Modern Electroplating*, Lowenheim, FA, ed., Wiley, New York, NY (1974), Chapter 31). However, such metal deposition processes must produce structures with excellent metal-to-metal adhesion and minimal surface contamination to provide a high density of active sites on the catalytic metal electrode surface. This high density of active sites is a necessary property for efficient redox conversion of electroactive species such as hydrogen peroxide.

[0123] In exemplary embodiments of the invention, a substrate layer is first coated with a thin-film conductive layer by electrode deposition, surface sputtering, or other suitable process steps. In one embodiment, this conductive layer may be configured as multiple thin-film conductive layers, such as an initial chromium-based layer suitable for chemical adhesion to a polyimide substrate, followed by a gold-based thin-film layer and a chromium-based thin-film layer formed sequentially. In alternative embodiments, other electrode layer configurations or materials may be used. The conductive layer is then covered with a selected photoresist coating according to conventional photolithography techniques, and a contact mask may be applied over the photoresist coating for suitable optical imaging. The contact mask typically contains one or more conductor trace patterns for appropriate exposure of the photoresist coating, followed by an etching step to retain multiple conductive sensor traces on the substrate layer. In an exemplary sensor configuration designed for use as a subcutaneous glucose sensor, each sensor trace may contain three parallel sensor elements corresponding to three separate electrodes, such as a working electrode, a counter electrode, and a reference electrode.

[0124] A portion of the sensor's conductive layer is typically covered by an insulating capping layer, which is typically a material such as a silicone polymer and / or polyimide. The insulating capping layer can be applied in any desired manner. In an exemplary procedure, the insulating capping layer is applied as a liquid layer over the sensor trace, after which the substrate is rotated to distribute the liquid material as a thin film over the sensor trace and extend the liquid material as a thin film beyond the edge margin of the sensor trace that seals against the substrate layer. This liquid material can then undergo one or more suitable radiation and / or chemical and / or thermal curing steps as known in the art. In alternative embodiments, the liquid material can be applied using spraying techniques or any other desired application method. Various insulating layer materials can be used, such as photoimageable epoxy acrylates, with illustrative materials including photoimageable polyimide, product number 7020, available from OCG Corporation, West Paterson, NJ.

[0125] As described above, optionally, after exposing the sensor tip through the opening, a suitable electrode chemical substance defining the distal electrode can be applied to the sensor tip. In an illustrative sensor embodiment with three electrodes used as a glucose sensor, an enzyme (typically glucose oxidase) is disposed within one of the openings, thus coating one of the sensor tips to define the working electrode. The same coating as the working electrode can be provided to one or two of the other electrodes. Alternatively, other suitable uncoated chemicals, such as other enzymes, can be provided to the other two electrodes, or chemicals can be provided to define the reference and counter electrodes of the electrochemical sensor.

[0126] Methods for producing the extremely thin enzyme coatings of the present invention include spin coating, dip and dry processes, low-shear spraying, inkjet printing, screen printing, etc. Because those skilled in the art can readily determine the thickness of enzyme coatings applied using processes in the art, they can easily identify methods capable of producing the extremely thin coatings of the present invention. Typically, such coatings are vapor-crosslinked after application. Surprisingly, the material properties of sensors produced by these processes exceed those of sensors with coatings produced by electrodeposition, including improved lifetime, linearity, regularity, and an improved signal-to-noise ratio. Furthermore, embodiments of the present invention utilizing glucose oxidase coatings formed by such processes are designed to recycle hydrogen peroxide and improve the biocompatibility profile of such sensors.

[0127] Sensors produced using processes such as spin coating also avoid other problems associated with electrodeposition, such as those related to the stress on the material placed on the sensor during the electrodeposition process. Specifically, the electrodeposition process has been observed to generate mechanical stress on the sensor, for example, mechanical stress generated by tension and / or compressive forces. In some cases, such mechanical stress can produce sensors with coatings that tend to crack or delaminate to some extent. This has not been observed in coatings applied to sensors using spin coating or other low-stress processes. Therefore, yet another embodiment of the invention is a method for avoiding cracking and / or delamination of coatings on sensors due to electrodeposition, the method comprising applying the coating via a spin coating process.

[0128] Method for using the analytical analyte sensor device of the present invention A related embodiment of the present invention is a method for sensing an analyte in a mammal, the method comprising implanting an analyte sensor embodiment disclosed herein into a mammal, and then sensing a change in current at a working electrode and correlating the current change with the presence of the analyte, thereby sensing the analyte. The analyte sensor may be anoly polarized such that the working electrode sensing the current change is the anode, or catholy polarized such that the working electrode sensing the current change is the cathode. In one such method, the analyte sensor device senses glucose in the mammal. In alternative embodiments, the analyte sensor device senses lactate, potassium, calcium, oxygen, pH, and / or any physiologically relevant analytes in the mammal.

[0129] Certain analyte sensors with the structures discussed above possess many desirable characteristics that allow for a variety of methods for sensing analytes in mammals. For example, in such methods, an analyte sensor device implanted in a mammal is used to sense analytes within the mammalian body for more than one month, two months, three months, four months, five months, or six months. Typically, such an analyte sensor device implanted in a mammal senses changes in current in response to the analyte within 15 minutes, 10 minutes, 5 minutes, or 2 minutes of contact with the sensor. In such methods, the sensor can be implanted in various locations within the mammalian body, such as both vascular and non-vascular spaces.

[0130] Example The following examples are provided to aid in understanding the invention; however, it should be understood that the invention is not limited to the specific materials or procedures described in the examples. All materials used in the examples are available from commercial sources.

[0131] Example 1: Synthesis and characterization of illustrative polyurea / polyurethane polymers using conventional methods: The disclosure provided herein, combined with information known in the art, confirms that functional linear polyurethane / polyurea polymers can be formulated from a number of formulations, such as those disclosed in the following U.S. Patents: 5,777,060; 5,882,494; 6,642,015; and PCT Publications WO 96 / 30431; WO 96 / 18115; WO 98 / 13685; and WO 98 / 17995, the contents of which are incorporated herein by reference. Some of these polymers provide formulations suitable for use as glucose-limiting membranes (GLMs).

[0132] The standard GLM formulations used to manufacture embodiments of the present invention include: 25 mol% polymethylsiloxane (PDMS), trimethylsilyl-terminated, 25-35 centistokes; 75 mol% polypropylene glycol diamine (Jeffamine 600, a polyethylene glycol amine with a molecular weight of approximately 600); and 50 mol% diisocyanate (e.g., 4,4'-diisocyanate).

[0133] Such standard GLM formulations and their synthesis methods are disclosed, for example, in U.S. Patents 6,642,015, 5,777,060, and 6,642,015.

[0134] Another formulation used in embodiments of the invention is referred to as a “semi-permeable GLM” because its glucose permeability is observed to be half that of the standard formulation described above. In the standard GLM, the Jeffamine / PDMS molar ratio is 3 / 1. In contrast, in the “semi-permeable GLM,” this ratio is changed to make Jeffamine / PDMS = 12 / 1. This semi-permeable GLM can be used, for example, to reduce the weight percentage of GLM urea in the entire polymer blend to achieve a specific Isig (or glucose permeability). Moreover, the presence of more GLM-acrylate polymer in the polymer blend can enhance the adhesion between the polycarbonate polymer film layer and the proximal layer in the sensor (e.g., the layer containing glucose oxidase).

Claims

1. A method for increasing the thermal stability of a biocompatible membrane formed from a reaction mixture, said reaction mixture comprising: Diisocyanate; A hydrophilic polymer, wherein the hydrophilic polymer includes a hydrophilic diol or a hydrophilic diamine; Siloxane, wherein the siloxane has an amino, hydroxyl, or carboxylic acid functional group at its end; as well as Tin catalyst; The method includes forming the reaction mixture such that the catalyst is present in the reaction mixture in an amount of 0.05% to 0.2% w / w of the reaction mixture components; This increases the thermal stability of the biocompatible membrane compared to a comparable membrane formed from a reaction mixture in which the catalyst is present in the reaction mixture in an amount greater than 0.2% w / w of the reaction mixture. The reaction mixture further comprises polycarbonate diol; and, (a) The diisocyanate comprises: Hexamethylene diisocyanate in a weight percentage of 17% to 23%; Methylene diphenyl diisocyanate in weight percentages of 0% to 8.5%; (b) The hydrophilic diol or hydrophilic diamine comprises 28% to 51% by weight of JEFFAMINE 600 and / or JEFFAMINE 900; (c) The siloxane comprises 14% to 48% by weight of polydimethylsiloxane-A15, having an amino group at its end; (d) The polycarbonate diol comprises 7.5% to 19% by weight of poly(1,6-hexyl carbonate) diol.

2. The method of claim 1, wherein the catalyst is present in the reaction mixture in an amount less than 0.11% w / w of the reaction mixture.

3. The method of claim 1, wherein the catalyst is present in the reaction mixture in an amount of 0.1% to 0.2% w / w of the reaction mixture.

4. The method according to claim 1, wherein water is added as a chain extender in the reaction mixture of the polyurea-polyurethane copolymer.

5. The method according to claim 1, wherein the thermal stability of the biocompatible membrane is measured by observing the change in the molecular weight of the biocompatible membrane during maintenance at a temperature of 60°C for at least 3, 5, or 7 days.

6. An amperometric detection analyte sensor, comprising: basal layer; A conductive layer disposed on the substrate layer and including a working electrode; An analyte sensing layer disposed on the conductive layer; as well as An analyte modulation layer is disposed on the analyte sensing layer, wherein the analyte modulation layer: (a) Formed from a reaction mixture comprising: Diisocyanate; A hydrophilic polymer, wherein the hydrophilic polymer includes a hydrophilic diol or a hydrophilic diamine; Siloxane, wherein the siloxane has an amino, hydroxyl, or carboxylic acid functional group at its end; as well as A tin catalyst, wherein the catalyst is present in the reaction mixture in an amount of 0.05% to 0.2% w / w of the reaction mixture components; and (b) Exhibits greater thermal stability compared to an equivalent analyte modulation layer formed by the reaction mixture in which the catalyst is present in the formulation in an amount greater than 0.2% w / w of the reaction mixture; The reaction mixture further comprises polycarbonate diol; and, (c) The diisocyanate comprises: Hexamethylene diisocyanate in a weight percentage of 17% to 23%; Methylene diphenyl diisocyanate in weight percentages of 0% to 8.5%; (d) The hydrophilic diol or hydrophilic diamine comprises 28% to 51% by weight of JEFFAMINE 600 and / or JEFFAMINE 900; (e) The siloxane comprises 14% to 48% by weight of polydimethylsiloxane-A15, having an amino group at its end; (f) The polycarbonate diol comprises 7.5% to 19% by weight of poly(1,6-hexyl carbonate) diol.

7. The analyte sensor according to claim 6, further comprising at least one of the following: A protein layer, wherein the protein layer is disposed on the analyte sensing layer; or A capping layer disposed on an analyte sensor device, wherein the capping layer includes pores positioned on the capping layer to facilitate the analyte present in the in vivo environment from contacting and diffusing through the analyte modulation layer and contacting the analyte sensing layer.

8. The analyte sensor of claim 6, wherein the conductive layer comprises a plurality of electrodes, the plurality of electrodes comprising a working electrode, a counter electrode, and a reference electrode.

9. The analyte sensor of claim 8, wherein the conductive layer comprises a plurality of working electrodes and / or counter electrodes and / or reference electrodes; and optionally, the plurality of working electrodes, counter electrodes and reference electrodes are grouped together as units and distributed in position on the conductive layer in a unit repeating pattern.

10. A method for manufacturing an analyte sensor for implantation in mammals, the method comprising the steps of: Provide a base layer; A conductive layer is formed on the substrate layer, wherein the conductive layer contains a working electrode; An analyte sensing layer is formed on the conductive layer, wherein the analyte sensing layer contains an oxidoreductase; An analyte modulation layer is formed on the analyte sensing layer, wherein the analyte modulation layer (a) Formed from a reaction mixture comprising: Diisocyanate; A hydrophilic polymer, wherein the hydrophilic polymer includes a hydrophilic diol or a hydrophilic diamine; Siloxane, wherein the siloxane has an amino, hydroxyl, or carboxylic acid functional group at its end; as well as A tin catalyst, wherein the catalyst is present in the reaction mixture in an amount of 0.05% to 0.2% w / w of the reaction mixture components; and (b) Exhibits greater thermal stability compared to a comparable analyte modulating layer formed from the reaction mixture in which the catalyst is present in the formulation at an amount greater than 0.2% w / w of the reaction mixture. The reaction mixture further comprises polycarbonate diol; and: (c) The diisocyanate comprises: Hexamethylene diisocyanate in a weight percentage of 17% to 23%; Methylene diphenyl diisocyanate in weight percentages of 0% to 8.5%; (d) The hydrophilic diol or hydrophilic diamine comprises 28% to 51% by weight of JEFFAMINE 600 and / or JEFFAMINE 900; (e) The siloxane comprises 14% to 48% by weight of polydimethylsiloxane-A15, having an amino group at its end; (f) The polycarbonate diol comprises 7.5% to 19% by weight of poly(1,6-hexyl carbonate) diol.

11. The method of claim 10, wherein water is added as a chain extender in the reaction mixture of the polyurea-polyurethane copolymer.

12. The method of claim 10, wherein the catalyst is present in the reaction mixture in an amount less than 0.11% w / w of the reaction mixture.

13. The method of claim 10, wherein the catalyst is present in the reaction mixture in an amount of 0.1% to 0.2% w / w of the reaction mixture.

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