CN112088217A - Thermostable glucose limiting membrane for glucose sensors - Google Patents

Abstract

Embodiments of the 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 limited amount of catalyst and/or polycarbonate compound to provide such membranes with improved material properties, such as enhanced thermal and hydrolytic stability.

Description

Thermostable glucose limiting membranes for glucose sensors

Cross-reference to related applications

This application claims priority under Section 120 of US Patent Application Serial No. 15/981,681, filed on May 16, 2018, entitled "THERMALLY STABLE GLUCOSE LIMITING MEMBRANE FOR GLUCOSE SENSORS" , the contents of said US patent application are incorporated herein by reference.

technical field

The present invention relates to biosensors, such as glucose sensors for diabetes management, and materials for making such sensors, eg, polymeric compositions useful in biosensor membranes.

Background technique

Analyte sensors, such as biosensors, comprise devices that use biological elements to convert chemical analytes in a matrix into a detectable signal. There are many types of biosensors 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.

A typical glucose sensor works according to the following chemical reactions:

H ₂ O ₂ → O ₂ +2H ⁺ +2e ^- Equation 2

Glucose oxidase is used to catalyze the reaction between glucose and oxygen to produce gluconic acid and hydrogen peroxide as shown in Equation 1. _The H2O2 reacted electrochemically as shown in Equation ₂ , and the current was measured by a potentiostat. The stoichiometry of the reaction poses challenges for developing sensors in vivo. Specifically, for optimal sensor performance, the sensor signal output should be determined only by the analyte of interest (glucose) and not by any co-substrate ( _O2 ) or kinetically controlled parameters such as diffusion. If oxygen and glucose are present in equimolar concentrations, _H2O2 is stoichiometrically related to the amount of glucose reacted at the enzyme _; and the associated current that generates the sensor signal is proportional to the amount of glucose reacted with the enzyme. However, if there is not enough oxygen to react all the glucose with the enzyme, the current will be proportional to the oxygen concentration instead of the glucose concentration. Therefore, in order for the sensor to provide a signal that depends only on the glucose concentration, the glucose must be the limiting reagent, i.e. the _O concentration must exceed all potential glucose concentrations. The problem with using such glucose sensors in vivo, however, is that the oxygen concentration relative to glucose is low when the sensor is implanted in the body, a phenomenon that can compromise the accuracy of the sensor readings.

There are many methods for addressing hypoxia. One approach is to use a homogeneous polymer membrane with hydrophobic and hydrophilic regions to control oxygen and glucose permeability. For example, Van Antwerp et al. developed linear polyurea membranes comprising a combination of polyethylene glycol and silicone hydrophobic components that allow high oxygen permeability and hydrophilic components that allow limited glucose permeability (see, eg, U.S. Pat. 5,777,060, 5,882,494 and 6,642,015). While possessing many useful and desirable properties, over time such polymeric compositions may experience some degradation under high temperature and high humidity conditions. In view of this, there is a need in the art for more robust polymeric membrane compositions that can be used, for example, to address the hypoxia problem observed in glucose sensors incorporating glucose oxidase.

SUMMARY OF THE INVENTION

Embodiments of the present invention provide compositions useful in analyte sensors and methods for making and using such compositions and sensors. In an exemplary embodiment of the invention, the sensor is a glucose sensor comprising an analyte-modulated membrane formed from a polymerized reaction mixture containing a limited amount of catalyst to provide improvements to such membranes material properties such as enhanced thermal and hydrolytic stability. As disclosed herein, when these polymer compositions are used to form analyte-limiting membranes in glucose sensors, the resulting sensors exhibit enhanced performance compared to conventional polymer compositions formed from reaction mixtures using conventional amounts of catalysts Long-term stability profile.

The invention disclosed herein has many embodiments. One embodiment of the present invention is a method of increasing the thermal stability of a biocompatible film formed from a reaction mixture, the reaction mixture comprising: a diisocyanate; a hydrophilic polymer, the hydrophilic polymer comprising a hydrophilic Hydrophilic diols or hydrophilic diamines; siloxanes having amino, hydroxyl or carboxylic acid functional groups at the ends; and catalysts. 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% (eg, 0.1%) of the components of the reaction mixture, thereby being different from the catalyst in which the catalyst is present in an amount greater than 0.2% (eg, 0.1%). The reaction mixture present in the formulation in an amount equal to or equal to 0.2% of the reaction mixture increases the thermal stability of the biocompatible film compared to a comparable film formed by the reaction mixture. Optionally, the reaction mixture further includes additional components such as polycarbonate diols.

Another embodiment of the present invention is an 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 , the analyte sensing layer is disposed on the conductive layer; and an analyte modulating layer is disposed on the analyte sensing layer. In this embodiment, the analyte modulating layer is formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic polymer diamines; siloxanes having amino, hydroxyl or carboxylic acid functional groups at the termini; and catalysts. In this example, the catalyst is present in the reaction mixture in an amount less than 0.2% of the components of the reaction mixture such that the analyte modulating layer is associated with the catalyst in which the catalyst is present in an amount greater than or equal to 0.2% of the reaction mixture The amount of the reaction mixture present in the formulation exhibits greater thermal stability than comparable analyte modulating layers.

Yet another embodiment of the present invention is a method of making an analyte sensor for implantation in a mammal. This method embodiment includes the steps of: providing a base layer; forming a conductive layer on the base layer, wherein the conductive layer includes a working electrode; forming an analyte sensing layer on the conductive layer, wherein the analyte senses The sensing layer comprises an oxidoreductase; and an analyte modulating layer is then formed on the analyte sensing layer. In this embodiment, the analyte modulating layer is formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic polymer a siloxane having amino, hydroxyl, or carboxylic acid functional groups at the termini; and a catalyst present in an amount less than 0.2% (eg, 0.1%) of the components of the reaction mixture in the reaction mixture such that the reaction mixture exhibits a comparable analyte modulating layer formed from the reaction mixture wherein the catalyst is present in the formulation in an amount greater than or equal to 0.2% of the reaction mixture Greater thermal stability. Optionally, the reaction mixture further includes additional components such as polycarbonate diols.

Other objects, features and advantages of the present invention will become apparent to those skilled in the art from the following detailed description. It is to be understood, however, that the detailed description and specific examples, while indicating some embodiments of the invention, are given by way of illustration and not of limitation. Many changes and modifications may be made within the scope of the present invention without departing from its spirit, and the present invention includes all such modifications.

Description of drawings

Reference is now made to the drawings, in which like reference numerals refer to corresponding parts throughout the following figures:

1 provides a diagrammatic view of one embodiment of an amperometric detection analyte sensor with multiple layered materials/elements in accordance with one or more embodiments of the present invention;

2A-C illustrate chemical structures of raw materials used in polycarbonate urea glucose limiting membranes (GLMs) in accordance with one or more embodiments of the present invention. Figure 2A shows the chemical structures of PDMS and Jeffamine. Figure 2B shows 4,4'-methylenebis(cyclohexyl isocyanate) or HMDI and 4,4'-methylenebis(phenyl isocyanate) or MDI. Figure 2C shows the chemical structure of polycarbonate diol;

Figure 3 illustrates a GLM synthesis reaction according to one or more embodiments of the present invention;

4A-B illustrate a comparison of the morphology of various counter electrodes to the morphology of the working electrode after testing in accordance with one or more embodiments of the present invention. Figure 10A shows bubbles (or pits) created at the counter electrode after use. The formation of air bubbles at the counter electrode may trigger delamination or unwanted biological responses (due to texture changes or rough surfaces). Figure 10B shows that MDI_polycarbonate_GLM can enhance GLM adhesion so that no bubbles (or pits) are generated at the counter electrode after use;

Figure 5 shows a comparison of in vitro SITS data between standard 2xGLM and PCU_GLM (polycarbonate urea glucose limiting membrane) in accordance with one or more embodiments of the present invention;

6 illustrates E3 sensor morphology after 7-day SITS testing of standard 2xGLM-coated sensors and PCU_GLM-coated sensors in accordance with one or more embodiments of the present invention;

7A-B provide graphs of data from thermal degradation studies of various formulations and compositions of formulations. Figure 7A provides results from a thermal degradation study at 45°C comparing the degradation (as observed by molecular weight reduction) of polymeric materials useful as biocompatible membranes (eg, analyte modulating layers in glucose sensors) results, while Figure 7B provides results from a similar thermal degradation study conducted at 60°C. In Figures 7A and 7B, the "control" material is formed from the polymerization mixture in which the catalyst is present in the formulation in an amount greater than or equal to 0.2% of the polymerization mixture, and the "new" material is formed from the polymerization mixture in which the catalyst is present in the formulation It is present in the formulation in an amount less than 0.2% of the reaction mixture (0.1% in this case).

8A-E show results from thermal degradation studies of various formulations and compositions of formulations in accordance with one or more embodiments of the present invention. Figure 8A shows results from thermal degradation studies. Figures 8B-D show the composition of various formulations. Figure 8E provides data showing that the glucose permeability (Pg) of polycarbonate-GLM did not decrease after baking;

9 shows a summary of results from thermal/hydrolysis studies of various sample formulations in accordance with one or more embodiments of the present invention. Thermal degradation test results demonstrate that MDI and polycarbonate chains can contribute to (slow down) the GLM degradation process.

Detailed ways

Unless otherwise defined, all technical terms, symbols and other scientific or technical terms used herein are intended to have the meaning commonly understood by one of ordinary skill in the art to which this invention belongs. In some cases, for the sake of clarity and/or ease of reference, terms with commonly understood meanings are defined herein, and the inclusion of these definitions herein should not necessarily be construed as representing a material that is material to the meanings commonly understood in the art difference. Many of the techniques and procedures described or referenced herein are well understood and routinely employed by those skilled in the art using routine methods. Procedures involving the use of commercially available kits and reagents are generally performed according to protocols and/or parameters defined by the manufacturer, unless otherwise indicated. A number of terms are defined below. All publications mentioned herein are incorporated by reference to disclose and describe the methods and/or materials in connection with which the publications are cited. Publications cited herein are cited for their disclosure prior to the filing date of this application. Nothing herein should be construed as an admission that the inventor is not entitled to advance the publication date by virtue of an earlier priority date or priority invention date. Further, actual publication dates may differ from those shown and require independent verification.

It must be noted that, as used herein and in the appended claims, the singular forms "a/an" and "the" include plural referents unless the context clearly dictates otherwise. Thus, for example, reference to "an oxidoreductase" includes a variety of such oxidoreductases and equivalents thereof, and the like, known to those skilled in the art. All numbers recited in the specification and associated claims that refer to values that can be numerically characterized by values other than integers (eg, "50 mol %") should be understood to be modified by the term "about."

As used herein, the term "analyte" is a broad term and is used in its ordinary sense, including, but not limited to, referring to biological fluids (eg, blood, interstitial fluid, cerebrospinal fluid, lymph fluid, or urine) that can be Substances or chemical constituents in the fluid being analyzed. Analytes may comprise naturally occurring substances, man-made substances, metabolites and/or reaction products. In some embodiments, the analyte for measurement by the sensing region, device and method is glucose. However, other analytes are also contemplated, including but not limited to lactic acid. In certain embodiments, naturally occurring salts, sugars, proteins, fats, vitamins, and hormones in blood or interstitial fluid may constitute analytes. Analytes can be naturally present in biological fluids or endogenous; eg, metabolites, hormones, antigens, antibodies, and the like. Alternatively, the analyte may be introduced into the human body or be exogenous, eg, contrast agents for imaging, radioisotopes, chemical agents, fluorocarbon-based synthetic blood, or drugs or pharmaceutical compositions, including but not limited to insulin . Metabolites of drugs and pharmaceutical compositions are also expected analytes.

As used herein, the term "sensor" is a broad term and is used in its ordinary sense, including, but not limited to, one or more portions of an analyte monitoring device that detect an analyte. 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 the sensor body and fixed within the sensor body so as to be on the body An electrochemically reactive surface is formed at one location of the body, an electrical connection is formed at another location of the body, and a membrane system is formed that adheres to the body and covers the electrochemically reactive surface. During normal operation of the sensor, a biological sample (eg, blood or interstitial fluid) or a portion thereof is contacted with an enzyme (eg, glucose oxidase) (either directly or after passing through one or more membranes or domains); the biological sample (eg, glucose oxidase) or a portion thereof) resulting in the formation of a reaction product that allows the determination of analyte levels in a biological sample.

As discussed in detail below, embodiments of the present invention relate to the use of electrochemical sensors that exhibit a range of novel material and functional elements. Such sensors incorporate novel polymeric compositions to form robust analyte-modulated membranes with a unique set of technically desirable material properties, such as increased thermal stability. Electrochemical sensors of the present invention are designed to measure an analyte of interest (eg, glucose) or to indicate the concentration of an analyte or the concentration of a substance present in a fluid. 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 indicative of the concentration of the analyte of interest. Typically, sensors are of the type that sense the products or reactants of an enzymatic reaction between an analyte and an enzyme in the presence of oxygen as a measure of the analyte in vivo or in vitro. Such sensors include a polymeric membrane surrounding an enzyme through which the analyte migrates before reacting with the enzyme. The product is then measured using electrochemical methods, and thus the output of the electrode system serves as a measure of the analyte. In some embodiments, the sensor may use amperometric, coulometric, conductometric, and/or potentiometric titration techniques to measure the analyte.

Analyte-modulating compositions (eg, analyte-modulating compositions useful as glucose limiting membranes in amperometric glucose sensors) comprise polymeric compositions formed from biocompatible polyurea materials (see, eg, Biocompatible Polyurea The case of the material is incorporated by reference). Such compositions can exhibit stable glucose and oxygen permeability, low protein adsorption rates, and biocompatibility. However, due to the content of PEG chains, certain degradation problems can occur under high temperature conditions and/or high humidity conditions. As disclosed in detail below, we have discovered that certain carbonate and aromatic isocyanate compounds can be added to the polymerization to replace some portion of the PDMS and HMDI polymeric chain elements. Two compounds have been found to increase the thermal and hydrolysis resistance of these polymers under high temperature conditions and high humidity conditions. Additionally, the chemical structures of the compounds provide evidence that such compositions have very good electron beam resistance. Carbonate materials useful in embodiments of the present invention include, but are not limited to, polycarbonate diols (eg, butanediol or hexanediol or similar compounds). In an illustrative embodiment of the invention, the Mw of the carbonate material is 500 to 2000 Daltons. Aromatic isocyanate materials useful in embodiments of the present invention include, but are not limited to, MDI or similar compounds.

In addition to limited amounts of catalyst in the polymerization mixture, the addition of MDI can improve the thermal and e-beam resistance of polymeric compositions used as analyte modulating (eg glucose confinement) compositions through its benzene ring structure. The benzene ring also acts as a good free radical scavenger for preventing oxidation of polymer components. Polycarbonate diols can provide better heat resistance and hydrolysis resistance through their carbonate structure (relative to ether or ester chains). The addition of polycarbonate segments to the polymer backbone can prevent unwanted deformation of the polymer composition layer disposed on the electrodes of the amperometric glucose sensor. Both gas and water are generated on the counter electrode between the analyte-sensing layer (eg, an analyte-sensing layer comprising enzymes such as GOX) and the analyte-molding (eg, glucose-limiting membrane) layer, which occurs over long periods of time Sensor failure (signal drift) may result after use. In this case, the polycarbonate fragments in the GLM backbone can prevent/reduce the chain rotation of PDMS in the GLM membrane, so the glucose permeability (Pg) of the GLM does not gradually decrease over time because the hydrophilic chains (Jeffamine or PEG) is encapsulated/incorporated by hydrophobic PDMS chains, especially for low Pg GLMs. In certain embodiments, in order to make a homogeneous urethane/urea copolymer, the synthesis will involve 3 injections of starting material after different timings. Raw materials were injected in a 4-2-4 ratio at time = 0, 4, 12 hours, respectively. The addition of polycarbonate chains to GLMs prevents Pg changes/reductions due to rotation/entanglement of PDMS chains over time, especially for low-Pg GLM films. To reduce thermal/radiative/oxidative degradation, the desired MDI content in the final polymer may be 2% to 25%. In order to prevent film deformation or Pg reduction due to the rotation of the silicone chains over time, the desired polycarbonate content in the final polymer may be 8% to 30%. Polycarbonate GLM showed good adhesion to AP with no more pits (bubbles) formed after testing.

Embodiments of the invention disclosed herein provide sensor types for use in, for example, 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 use some form of immobilized enzyme to achieve their biological specificity. Embodiments of the invention described herein can be adapted and implemented using a variety of known electrochemical sensors including, for example, US Patent Application Nos. 20050115832, US Patent Nos. 6,001,067, 6,702,857,第6,212,416号、第6,119,028号、第6,400,974号、第6,595,919号、第6,141,573号、第6,122,536号、第6,512,939号、第5,605,152号、第4,431,004号、第4,703,756号、第6,514,718号、第5,985,129号、第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 Publications 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 which are incorporated herein by reference.

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 (eg, including sensors and, eg, monitors, sensor systems of associated electronic components such as processors). The present disclosure further provides methods for making and using such sensor and/or architectural configurations. While some embodiments of the present invention relate to glucose sensors and/or lactate sensors, the various elements disclosed herein (eg, analyte modulating membranes made from polycarbonate polymeric compositions) can be adapted for use with various components known in the art. Use with any of a wide variety of sensors. The analyte sensor elements, architectures, and methods for making and using these elements disclosed herein can be used to create various 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 properties of a wide variety of analyte species.

Specific aspects of embodiments of the invention are discussed in detail in the following sections.

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 comprising amperometric sensors for the detection and/or measurement of biological analytes such as glucose are known in the art. Many glucose sensors are based on oxygen (Clark-type) amperometric detection transducers (see, eg, Yang et al., Electroanalysis, 1997, 9, No. 16: 1252-1256; Clark et al., Ann. NY, Academy of Sciences ( Acad. Sci.)" 1962, 102, 29; Updike et al., "Nature" 1967, 214, 986; and Wilkins et al., "Med. Engineering. 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 easily 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 the use of one or more polycarbonate polymeric films disclosed herein that can modulate the transport properties of different compounds whose reactions are detected in hydrogen peroxide-based amperometric detection A signal is generated at the transducer element. Thus, these membranes can be used with, for example, various H2O2 _- based analyte _sensors that benefit from optimized reaction stoichiometry.

As noted above, embodiments of the present invention include sensor membranes made from reaction mixtures to include limited amounts of catalyst and/or polycarbonate polymer compositions. As known in the art, polymers include long or larger molecules consisting of chains or networks of many repeating units formed by chemically bonding many identical or similar small molecules called monomers in formed together. A copolymer or heteropolymer is a polymer derived from two (or more) monomeric species, as opposed to a homopolymer that uses 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 polymeric compositions disclosed herein can optimize analyte sensor function including sensor sensitivity, stability, and hydration profile. Additionally, by optimizing the stoichiometry of reactant species over the sensor temperature range, the membranes disclosed herein can optimize chemical reactions that generate key measurable signals related to analyte (eg, glucose) levels of interest. The following sections describe illustrative sensor elements, sensor configurations, and method embodiments of the present invention.

The polymeric materials disclosed herein are useful in various contexts as biocompatible membranes, such as glucose limiting membranes (GLMs). However, polymers often degrade over time due to oxidation, UV light, heat, hydrolysis or other processes. In this case, it has been found that traces of tin catalyst residues in polymer mixtures used to form biocompatible films such as GLM can accelerate GLM degradation over time. For this reason, older sensors may perform slightly worse than newly fabricated sensors due to the gradual degradation of the GLM. While not bound by a particular scientific theory or mechanism of action, it is believed that trace amounts of tin catalyst residues in GLMs may further trigger immune responses and loss of sensitivity. In this case, it was further found that reducing the amount of tin catalyst used in the GLM synthesis process (eg a 50% reduction) unexpectedly yielded films with increased resistance to thermal degradation, and improved the quality (biocompatibility) of the GLM and thermal stability) and glucose sensor in vivo performance. This greater GLM stability further extends the shelf life of the sensor and improves the biocompatibility of the sensor without any significant manufacturing process changes.

The invention disclosed herein has many embodiments. One embodiment of the present invention is a method of increasing the thermal stability of a biocompatible film formed from a reaction mixture, the reaction mixture comprising: a diisocyanate; a hydrophilic polymer, the hydrophilic polymer comprising a hydrophilic Hydrophilic diols or hydrophilic diamines; siloxanes having amino, hydroxyl or carboxylic acid functional groups at the ends; and catalysts. 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% (eg, 0.1%) of the components of the reaction mixture, thereby being different from the catalyst in which the catalyst is present in an amount greater than 0.2% (eg, 0.1%). The reaction mixture present in the formulation in an amount equal to or equal to 0.2% of the reaction mixture increases the thermal stability of the biocompatible film compared to a comparable film formed by the reaction mixture. Optionally, the reaction mixture (eg, at 60°C) uses an organic solvent such as tetrahydrofuran and further includes additional components such as polycarbonate diol. The thermal stability of various biocompatible films fabricated in this manner can be measured by various art-accepted practices, such as by observing maintenance at a temperature of 60°C for a period of at least 3, 5, or 7 days The change in molecular weight of the biocompatible membrane (see e.g. Figure 7) was measured.

In certain embodiments of the invention, the tin catalyst (eg, dibutyltin bis(ethyl 2-hexanoate)) is present in an amount less than 0.19%, 0.17%, 0.15%, 0.13%, or 0.11% of the reaction mixture (eg, 0.1 %) present in the reaction mixture. In some embodiments of the present invention, the diisocyanates include hexamethylene diisocyanate and/or methylene diphenyl diisocyanate, and/or hydrophilic diols including hydrophilic diols or hydrophilic diamines Sex polymers include JEFFAMINE, and/or siloxanes with amino, hydroxyl, or carboxylic acid functional groups at the termini include polydimethylsiloxane, and/or polycarbonate diols include poly(1,6-hexanecarbonate) ester) diol and/or poly(1,6hexyl-1,5amyl carbonate)diol. For example, in certain embodiments of the present 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 includes 28% to 51% by weight of JEFFAMINE 600 and/or JEFFAMINE 900, and the polydimethylsiloxane includes 14% to 48% by weight of Dimethicone-A15, and The polycarbonate diol includes 7.5% to 19% by weight of poly(1,6-hexyl carbonate) diol. In a specific illustrative embodiment, the diisocyanate includes about 22% hexamethylene diisocyanate and about 3.5% methylene diphenyl diisocyanate, JEFFAMINE includes about 45% JEFFAMINE 600 and/or JEFFAMINE 900, poly The dimethylsiloxane included about 22.5% polydimethylsiloxane-A15, and the polycarbonate diol included about 7.5% poly(1,6-hexyl carbonate) glycol. Optionally, in this method, water is added as a chain extender to the polyurea-polyurethane copolymer reaction mixture. Illustrative reaction mixtures of the present invention are shown in Tables 1 and 2 below.

Table 1

Table 2

Another embodiment of the present invention is an 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 , the analyte sensing layer is disposed on the conductive layer; and an analyte modulating layer is disposed on the analyte sensing layer. In this embodiment, the analyte modulating layer is formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic polymer diamines; siloxanes having amino, hydroxyl or carboxylic acid functional groups at the termini; and catalysts. In this example, the catalyst is present in the reaction mixture in an amount less than 0.2% of the components of the reaction mixture such that the analyte modulating layer is associated with the catalyst in which the catalyst is present in an amount greater than or equal to 0.2% of the reaction mixture The amount of the reaction mixture present in the formulation exhibits greater thermal stability than comparable analyte modulating layers. Optionally, the reaction mixture further includes additional components such as polycarbonate diols.

In typical embodiments, the analyte sensor is a glucose sensor that can be implanted in vivo. Optionally, the analyte sensor further comprises at least one of: a protein layer disposed over the analyte sensing layer; or a cover layer disposed over the analyte on the sensor device, and the cover layer includes holes positioned on the cover layer to facilitate analytes present in the in vivo environment from contacting and diffusing through the analyte modulating layer; and contacting the analysis object sensing layer. In some of these analyte sensors, the conductive layer includes a plurality of electrodes including a working electrode, a counter electrode, and a reference electrode, such as one embodiment wherein the conductive layer includes a plurality of electrodes and optionally, the plurality of working electrodes, counter electrodes and reference electrodes are grouped together as a unit and are geographically distributed in the unit repeating pattern on the conductive layer.

Yet another embodiment of the present invention is a method of making an analyte sensor for implantation in a mammal. This method embodiment includes the steps of: providing a base layer; forming a conductive layer on the base layer, wherein the conductive layer includes a working electrode; forming an analyte sensing layer on the conductive layer, wherein the analyte senses The sensing layer comprises an oxidoreductase; and an analyte modulating layer is then formed on the analyte sensing layer. In this embodiment, the analyte modulating layer is formed from a reaction mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic polymer a siloxane having amino, hydroxyl, or carboxylic acid functional groups at the termini; and a catalyst present in an amount less than 0.2% (eg, 0.1%) of the components of the reaction mixture in the reaction mixture such that the reaction mixture exhibits a comparable analyte modulating layer formed from the reaction mixture wherein the catalyst is present in the formulation in an amount greater than or equal to 0.2% of the reaction mixture Greater thermal stability. Optionally, the reaction mixture further includes additional components such as polycarbonate diols.

In a certain method of making an analyte sensor for implantation in a mammal, the diisocyanate comprises hexamethylene diisocyanate and/or methylene diphenyl diisocyanate, JEFFAMINE comprises about 45% JEFFAMINE 600 and /or JEFFAMINE 900, the polydimethylsiloxane includes about 22.5% polydimethylsiloxane-A15, and the polycarbonate diol includes about 7.5% poly(1,6-hexyl carbonate) glycol. Typically, in this example, the catalyst (eg, dibutyltin bis(ethyl 2-hexanoate)) is present in an amount less than 0.19%, 0.17%, 0.15%, 0.13%, or 0.11% of the reaction mixture (eg, about 0.1 %) present in the reaction mixture.

A certain amperometric detection sensor design for use with embodiments of the present invention includes multiple layered elements including, for example, a substrate layer with electrodes, an analyte sensing layer (eg, an analyte sensing layer including glucose oxidase), and an analytical An analyte modulating layer that functions in analyte diffusion control (eg, to modulate the amount of glucose and oxygen exposed to the analyte sensing layer). One such sensor embodiment is shown in FIG. 1 . Layered sensor designs incorporating the polycarbonate polymeric compositions disclosed herein as analyte modulating layers exhibit a range of material properties that overcome observations observed in various sensors including in vivo implanted electrochemical glucose sensors challenge. For example, sensors designed to measure analytes in aqueous environments (eg, sensors implanted in vivo) typically require the layers to be wetted before and during accurate analyte readings. Because the properties of a material may affect its rate of hydration, the material properties of a membrane used in an aqueous environment would ideally facilitate sensor wetting, such as to allow the sensor to be introduced into an aqueous environment and the sensor to provide analyte concentrations that are related to the environment. The time period between the corresponding accurate signal capabilities. Embodiments of the present invention including polycarbonate polymeric compositions address such issues by promoting sensor hydration.

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 modulating layer should not exacerbate (and ideally should decrease) as is known in the art For the case of "hypoxia problem". Specifically, because glucose oxidase-based sensors require both oxygen (O ₂ ) and glucose to generate a signal, an excess of oxygen relative to excess of glucose is necessary for the operation of a glucose oxidase-based glucose sensor. However, since the oxygen concentration in the subcutaneous tissue is much lower than the glucose concentration, oxygen may be the limiting reactant in the sensor in the reaction between glucose, oxygen and glucose oxidase, a situation that compromises the sensor production strictly dependent on The ability to signal glucose concentration. In this case, the chemical reaction between glucose and glucose oxidase is used to generate the measurable signal 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 modulating layer used in the electrochemical glucose sensor should not favor the diffusion of glucose over oxygen, for example, in a way that causes hypoxia problems. Embodiments of the present invention that include the polycarbonate polymeric compositions disclosed herein do not cause anoxic problems, but function to ameliorate them.

In addition, sensor designs using the polycarbonate polymeric compositions disclosed herein as analyte modulating layers can also overcome the use of analytes that can exhibit different diffusion profiles (eg, rates of analyte diffusion through the sensor material) at different temperatures. The complexity observed in the case of sensor materials. Specifically, for optimal sensor performance, the sensor signal output over a certain temperature range should be determined only by the level of the analyte of interest (eg glucose) and not by any co-substrate (eg O ₂ ) or kinetic control parameter decision. However, as is known in the art, the diffusion of compounds through the polymer matrix can be temperature dependent. Such temperature-dependent diffusion profiles may affect the reliance on which the sensor signal is generated in the event that an analyte (eg, glucose) diffuses through the polymer to react at a site where it reacts with another compound (eg, glucose oxidase). The stoichiometry of the reaction, whereby the technician makes the effort to make the sensor signal output depend only on the concentration of the analyte of interest over a range of temperatures. Accordingly, analyte-modulating compositions made from materials with analyte (eg, glucose) diffusion profiles that are stable over a range of temperatures (eg, 22 to 40 degrees Celsius) address such problems.

The invention disclosed herein provides polycarbonate polymeric compositions useful as membranes for biosensors such as, for example, amperometric glucose sensors. Embodiments of the present invention include, for example, a sensor having a plurality of layered elements comprising an analyte-limiting membrane comprising a polycarbonate polymeric composition. Such polymeric membranes are particularly useful in the construction of electrochemical sensors for in vivo use. Membrane embodiments of the present invention allow for a combination of desirable properties including enhanced hydration profile and stable permeability to molecules such as glucose over a range of temperatures. In addition, these polymeric films exhibit good mechanical properties for use as outer polymeric films. Therefore, glucose sensors incorporating such polymeric membranes show a very desirable in vivo performance profile.

Embodiments of the present invention include both materials (eg, polycarbonate polymeric compositions) as well as architectures designed to facilitate sensor performance. For example, in some embodiments of the present invention, the conductive layer includes a plurality of working and/or counter and/or reference electrodes (eg, 3 working electrodes, a reference electrode, and a counter electrode), so that For example, to avoid problems associated with poor sensor hydration and/or to provide redundant sensing capabilities. Optionally, the plurality of working, counter and reference electrodes are arranged together as a unit and are positionally distributed on the conductive layer in a repeating unit pattern. In certain embodiments of the invention, the base layer is made of a flexible material that allows the sensor to twist and bend when implanted in the body; and the electrodes are grouped into the following configurations: when the sensor device is implanted in the body The in-vivo fluid contacts at least one of the working electrodes when twisted and flexed. In some embodiments, the electrodes are grouped into a configuration that allows the sensor to continue to function if a portion of the sensor having one or more electrodes is removed from the in vivo environment and exposed to the ex vivo environment. Typically, the sensor is operably coupled to: a sensor input capable of receiving a signal from the sensor based on the sensed analyte; and a processor coupled to the sensor input, wherein the processor One or more signals received from a sensor can be characterized. In some embodiments of the invention, a pulsed voltage is used to obtain a signal from one or more electrodes of the sensor.

The sensors disclosed herein can be made from a wide variety of materials known in the art. In an illustrative embodiment of the invention, the analyte modulating layer comprises a polyurethane/polyurea polymer formed from a mixture comprising: a diisocyanate; a hydrophilic polymer comprising a hydrophilic polymer Aqueous diols or hydrophilic diamines; and siloxanes having amino, hydroxyl or carboxylic acid functional groups at the ends; with this polymer, polymerized with branched acrylates formed from the mixture Polycarbonates of these compounds include: butyl acrylate, propyl acrylate, ethyl acrylate, or methyl acrylate; amino acrylates; siloxane-acrylates; and poly(ethylene oxide)-acrylates. Optionally, additional materials may be included in these polymer blends. For example, certain embodiments of branched acrylate polymers are formed from reaction mixtures comprising hydroxy acrylate compounds such as 2-hydroxyethyl methacrylate.

The term "polyurethane/polyurea polymer" as used herein refers to polymers containing urethane linkages, urea linkages, or combinations thereof. As known in the art, polyurethanes are polymers composed of chains of organic units linked by urethane (urethane) chains. Polyurethane polymers are typically formed by step-growth polymerization by reacting a monomer containing at least two isocyanate functional groups with another monomer containing at least two hydroxyl groups (alcohols) in the presence of a catalyst. Polyurea polymers are derived from the reaction product of an isocyanate component and a diamine. Typically, such polymers are formed by combining diisocyanates with alcohols and/or amines. For example, combining isophorone diisocyanate with PEG 600 and aminopropylpolysiloxane under polymerization conditions provides a polyurethane/polyurea composition having both urethane (urethane) chains and urea linkages. Such polymers are known in the art and described in, for example, US Patent Nos. 5,777,060, 5,882,494 and 6,632,015, and PCT Publication Nos. WO 96/30431; WO 96/18115 ; WO 98/13685; and WO 98/17995, the contents of each of which are incorporated by reference.

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

Diisocyanates useful in this embodiment of the present invention are those typically used to prepare biocompatible polyurethanes. Such diisocyanates are described in detail in Szycher, SEMINAR ONADVANCES IN MEDICAL GRADE POLYURETHANES, Technomic Publishing, (1995) and include both aromatic diisocyanates and aliphatic diisocyanates. By. Examples of suitable aromatic diisocyanates include tolyl diisocyanate, 4,4'-methylene diphenyl diisocyanate, 3,3'-dimethyl-4,4'-biphenyl diisocyanate, Naphthalene diisocyanate and p-phenylene diisocyanate. Suitable aliphatic diisocyanates include, for example, 1,6 hexamethylene diisocyanate (HDI), trimethylhexyl diisocyanate (TMDI), trans-1,4-cyclohexane diisocyanate (CHDI), 1,4- Cyclohexane bis(methylene isocyanate) (BDI), 1,3-cyclohexane bis(methylene isocyanate) (H ₆ XDI), isophorone diisocyanate (IPDI) and 4,4'-idene Methylbis(cyclohexylisocyanate) ( _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, Wis., USA), or can be obtained by using literature Procedures are readily prepared by standard synthetic methods.

The amount of diisocyanate in the reaction mixture for the polyurethane/polyurea polymer composition is typically about 50 mole percent 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 to react 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 hydrophilic polymer (diol, diamine, or a combination thereof), b moles of silicone polymer with functionalized ends, and c moles of chain extender such that x=a+b+c, it should be understood that c may be zero.

Another reactant used to prepare the polyurethane/polyurea polymers 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) glycol, a polyester-based polyol, or a polycarbonate polyol. The term "poly(alkylene) glycol" as used herein refers to lower alkylene glycols such as poly(ethylene) glycol, poly(propylene) glycol, and polytetramethylene ether glycol (PTMEG) and the like Polymers of alkyl glycols. The term "polyester-based polyol" refers to a polymer in which the R group is a lower alkylene group such as ethylene, 1,3-propene, 1,2-propene, 1,4-butene , 2,2-dimethyl-1,3-propene, etc. (eg, as depicted in Figure 4 of US Pat. No. 5,777,060). Those skilled in the art will also appreciate that the diester portion of the polymer can also be different from the hexacarbondioic acid shown. For example, although Figure 4 of US Patent No. 5,777,060 shows an adipic acid component, the present invention also contemplates the use of succinates, glutarates, and the like. 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, longer chain aliphatic, cycloaliphatic, or aromatic groups. The term "hydrophilic diamine" refers to any of the aforementioned hydrophilic diols in which the terminal hydroxyl group has been substituted with a reactive amine group or in which the terminal hydroxyl group has been derivatized to yield a chain extension. For example, a certain hydrophilic diamine is a "diaminopoly(oxyalkylene)," which is a poly(alkylene)glycol in which the terminal hydroxyl group is replaced by an amino group. The term "diaminopoly(oxyalkylene)" also refers to poly(alkylene) glycols having aminoalkyl ether groups at the chain ends. An example of a suitable diamino poly(oxyalkylene) 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 polymer synthesis.

The amount of hydrophilic polymer used to prepare the linear polymer composition is generally from about 10 mole percent to about 80 mole percent relative to the diisocyanate used. Typically, the amount is from about 20 mole percent to about 60 mole percent relative to the diisocyanate. When lower amounts of hydrophilic polymers are used, it is common to include chain extenders.

The silicone-containing polyurethane/polyurea polymers useful in the present invention are generally linear, have excellent oxygen permeability and are substantially free of glucose permeability. Typically, the silicone polymer is a polydimethylsiloxane with two reactive functional groups (ie, a functionality of 2). The functional group can be, for example, a hydroxyl, amino or carboxylic acid group, but is usually a hydroxyl or amino group. In some embodiments, a combination of silicone polymers may be used, wherein the first moiety includes hydroxyl groups and the second moiety includes amino groups. Typically, functional groups are located at the chain ends of the silicone polymer. Many suitable silicone polymers are available from, for example, The Dow Chemical Company (Midland, Mich., USA) and General Electric Company (Ske, NY, USA) Commercially obtained from sources such as Silicones Division, Schenectady, N.Y., USA). Still other silicone polymers can be obtained from commercially available siloxanes (United Chemical Technologies, Bristol, PA, USA) by general synthetic methods known in the art (see, eg, US Pat. No. 5,777,060). Technologies, Bristol. Pa., USA). For use in the present invention, the silicone polymer is typically a silicone polymer having a molecular weight of about 400 to about 10,000, more typically a molecular weight of about 2000 to about 4000 The amount of silicone polymer incorporated into the reaction mixture will depend on the desired properties of the resulting polymer to form the biocompatible film. For compositions where lower glucose permeation is desired, larger amounts may be employed of silicone polymer. Alternatively, for compositions where higher glucose permeation is desired, a smaller amount of silicone polymer may be employed. Typically, for glucose sensors, the amount of silicone polymer is 10 moles relative to the diisocyanate % to 90 mol %. Typically, the amount is about 20 mol % to 60 mol % relative to the diisocyanate.

In one set of embodiments, the reaction mixture used to prepare the biocompatible film will also contain a chain extender which is an aliphatic or aromatic diol, aliphatic or aromatic diamine, alkane Alcohol amines or combinations thereof (eg, as depicted in Figure 8 of US 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, 1,4-cyclohexanediol Hexane dimethanol. Aromatic chain extenders include, for example, p-bis(2-hydroxyethoxy)benzene, m-bis(2-hydroxyethoxy)benzene, Ethacure 100(iii) (2,4-diamino-3,5-diethyltoluene) mixture of two isomers), Ethacure 300(iii) (2,4-diamino-3,5-bis(methylthio)toluene), 3,3'-dichloro-4,4'diaminodiphenylmethane, Polacure (iii) 740M (trimethylene glycol bis(p-aminobenzoic acid) ester) and diphenylaminomethane. Incorporation of one or more of the above chain extenders generally provides additional physical strength to the resulting biocompatible membrane without significantly increasing the glucose permeability of the polymer. Typically, chain extenders are used when lower amounts (ie, 10-40 mol%) of hydrophilic polymer are used. Specifically in some compositions, the chain extender is diethylene glycol present at about 40 to 60 mole % relative to the diisocyanate.

The polymerization of the above-mentioned reactants can be carried out in batch form or in a solvent system. The use of a catalyst is partially required, but not required. Suitable catalysts include dibutyltin bis(ethyl 2-hexanoate) (CAS number: 2781-10-4), dibutyltin diacetate, triethylamine, and combinations thereof. Typically, dibutyltin bis(ethyl 2-hexanoate) is used as a catalyst. Typical amounts of this catalyst used in the formulation are 0.05% to 0.2% (w/w ratio). Bulk polymerization is typically performed at an initial temperature of about 25°C (ambient temperature) to about 50°C to ensure adequate mixing of the reactants. As the reactants are mixed, an exotherm is usually observed, with the temperature rising to about 90-120°C. After the initial exotherm, the reaction flask can be heated at 75°C to 125°C, with 90°C to 100°C being an exemplary temperature range. Heating is usually performed for one to two hours. Solution polymerization can be performed in a similar manner. Solvents suitable for solution polymerization include dimethylformamide, dimethylsulfoxide, dimethylacetamide, halogenated solvents such as 1,2,3-trichloropropane, and ketones such as 4-methyl-2-pentanone . Typically, THF is used as a solvent. When the polymerization is carried out in a solvent, the heating of the reaction mixture is usually carried out for three to four hours.

The polymers prepared by bulk polymerization are usually dissolved in dimethylformamide and precipitated from water. Polymers prepared in water-immiscible solvents can be isolated by vacuum stripping of the solvent. These polymers were then dissolved in dimethylformamide and precipitated from water. After thorough washing with water, the polymer can be vacuum dried to constant weight at about 50°C.

Film preparation can be accomplished by dissolving the dried polymer in a suitable solvent and casting the film onto a glass plate. The choice of a suitable solvent for casting generally depends on the particular polymer and the volatility of the solvent. Typically, the solvent is _THF , _CHCl3 , _CH2Cl2 , DMF, IPA, or a combination thereof. More typically, the solvent is THF or DMF/ _CH2Cl2 ( ₂ /98% by volume). The solvent was removed from the film, the resulting film was fully hydrated, its thickness was measured and the water pickup was determined. Films useful in the present invention typically have a water absorption of about 20 to about 100 wt%, typically 30 to about 90 wt%, and more typically 40 to about 80 wt%.

Oxygen and glucose diffusion coefficients can also be determined for the individual polymer compositions as well as for the polycarbonate polymeric films of the present invention. Methods for determining diffusion coefficients are known to those skilled in the art, and examples are provided below. Certain embodiments of the biocompatible films described herein typically have an oxygen diffusion coefficient (D _oxygen ) of about 0.1 x 10" ^{6 cm2} /sec to about 2.0 x 10" ^{6 cm2} /sec and about 1 x 10 ^" Glucose diffusion coefficient (D _glucose ) of ⁹ cm ² /sec to about 500 x 10 ⁻⁹ cm ² /sec. More typically, the glucose diffusion coefficient is about 10 x 10" ^{9 cm2} /sec to about 200 x 10" ^{9 cm2} /sec.

Typical combinations of sensor elements

Embodiments of the present invention further comprise sensors comprising the polycarbonate polymeric compositions disclosed herein in combination with other sensor elements, such as interference suppression films (eg, as disclosed in US Patent Application Serial No. 12/572,087, The contents of said US patent application are incorporated by reference). One such embodiment of the present invention is an interference suppression film comprising a methacrylic acid polymer having a molecular weight between 100 kilodaltons and 1000 kilodaltons, wherein the methacrylic acid The polymers are crosslinked by hydrophilic crosslinking agents such as organofunctional dipodalalkoxysilanes. Another embodiment of the present invention is an interference suppressing membrane having a molecular weight between 4,000 Daltons and 500 kilodaltons of primary amine polymers, wherein the primary amine polymers are produced by, for example, pentanediol Aldehyde and other hydrophilic cross-linking agents are cross-linked. Typically, these interference suppressing films are coated with a hydrogen peroxide conversion composition. In an illustrative embodiment, the hydrogen peroxide conversion composition includes an electrode; and the crosslinked interference suppression film is coated on the electrode in a layer of 0.1 μm and 1.0 μm thick.

In some embodiments of the invention, elements of the sensor device, such as electrodes or apertures, are designed to have specific configurations and/or be made of specific materials and/or positioned relative to other elements in order to facilitate the function of the sensor. In one such embodiment of the invention, the working electrode, the counter electrode and the reference electrode are geographically distributed on the substrate layer and/or the conductive layer in the following configuration: when the sensor device is placed in contact with a fluid comprising the analyte (eg by inhibiting shading of the electrodes, which can inhibit hydration and capacitive activation of the sensor circuit) Facilitates sensor activation and/or maintains hydration of the working, counter, and/or reference electrodes. Generally, such embodiments of the present invention facilitate sensor startup and/or initialization.

Optionally, embodiments of the device include multiple working and/or counter and/or reference electrodes (eg, 3 working electrodes, one reference electrode, and one counter electrode), eg, to provide redundant sensing ability. Certain embodiments of the present 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 plurality of working, counter and reference electrodes are arranged together as a unit and are positionally distributed on the conductive layer in a repeating unit pattern. In certain embodiments of the present invention, the elongated base layer is made of a flexible material that allows the sensor to twist and bend when implanted in the body; and the electrodes are grouped into the following configurations: when the sensor device is in the The in vivo fluid contacts at least one of the working electrodes when twisted and bent when implanted in vivo. In some embodiments, the electrodes are grouped into a configuration that allows the sensor to continue to function if a portion of the sensor having one or more electrodes is removed from the in vivo environment and exposed to the ex vivo environment.

In certain embodiments of the present invention that include multiple sensors, elements such as sensor electrodes are organized/disposed within a flexible circuit assembly. In such embodiments of the invention, the architecture of the sensor system may be designed such that the first sensor does not affect the signal etc. generated by the second sensor (and vice versa); The tissue envelope of the sensor is used for sensing; so the signals from the individual sensors do not interact. Also, in typical embodiments of the invention, the sensors will be spaced a distance from each other, allowing the sensors to be easily packaged together and/or suitable for implantation by a single insertion action. One such embodiment of the present invention is a device for monitoring an analyte in a patient, the device comprising: a base element adapted to secure the device to a patient; a first piercing member, the the first piercing member is coupled to and extends from the base element; a first electrochemical sensor operably coupled to the first piercing member and including a first electrochemical sensor electrode, the first electrochemical sensor electrode for determining at least one physiological characteristic of the patient at the first electrochemical sensor placement site; a second piercing member coupled to and from the base element extension; a second electrochemical sensor operably coupled to the second piercing member and including a second electrochemical sensor electrode for use in the second electrochemical sensor At least one physiological characteristic of the patient is determined at the sensor placement site. In such embodiments of the invention, the at least one physiological property monitored by the first electrochemical sensor or the second electrochemical sensor includes the concentration of a naturally occurring analyte in the patient; the first piercing member connects the first electrochemical sensor to the Disposed in the first tissue compartment of the patient, and the second piercing member disposes the second electrochemical sensor in the second tissue compartment of the patient; and the first piercing member and the second piercing member are selected to avoid The configuration of the sensor signal generated by the second electrochemical sensor, possibly due to a physiological response change caused by implantation of the first electrochemical sensor, is disposed on the substrate.

Various elements of the sensor device may be positioned at certain locations in the device and/or configured in specific shapes and/or constructed from specific materials to facilitate the strength and/or function of the sensor. One embodiment of the present invention includes an elongated substrate comprising a polyimide or dielectric ceramic material that promotes the strength and durability of the sensor. In certain embodiments of the invention, the structural features and/or relative positions of the working electrode and/or the counter electrode and/or the reference electrode are designed to affect the fabrication, use and/or function of the sensor. Optionally, the sensor is operably coupled to a series of elements including a flexible circuit (eg, electrodes, electrical conduits, contact pads, etc.). One embodiment of the present invention includes an electrode having one or more rounded edges to inhibit delamination of a layer disposed on the electrode (eg, an analyte sensing layer including glucose oxidase).

In certain embodiments of the present invention, the electrodes of the device include a platinum composition, and the device further includes a titanium composition disposed between the elongated base layer and the conductive layer. Optionally, in such embodiments, the device further includes a gold composition disposed between the titanium composition and the conductive layer. Such embodiments of the present invention generally exhibit enhanced bonding and/or less corrosion and/or improved biocompatibility profiles between layered materials within the sensor. Related embodiments of the present invention include methods for inhibiting corrosion of sensor elements and/or methods for improving the biocompatibility of sensor embodiments of the present invention (eg, sensors constructed to use such materials).

In typical embodiments of the invention, sensors are operably coupled to additional elements (eg, electronic components), such as elements designed to transmit and/or receive signals, monitors, processors, etc., and sensor data may be used A device that modulates a patient's physiology, such as a drug infusion pump. For example, in some embodiments of the invention, a sensor is operably coupled to: a sensor input capable of receiving a signal from the sensor based on a value of a physiological characteristic sensed in the mammal; and a processor, so The processor is coupled to the sensor input, wherein the processor is capable of characterizing one or more signals received from the sensor. A wide variety of sensor configurations as disclosed herein can be used in such systems. Optionally, for example, the sensor includes three working electrodes, one counter electrode, and one reference electrode. In certain 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.

Diagrammatic view of typical sensor configuration

Figure 1 shows a cross-section of an exemplary sensor embodiment 100 of the present invention. This sensor embodiment is formed from a plurality of components, generally in the form of layers of various conductive and non-conductive compositions disposed on top of each other according to the art-recognized methods disclosed herein and/or the specific methods of the present invention. form. The components of the sensor are generally characterized herein as layers because, for example, it allows the sensor structure shown in FIG. 1 to be easily characterized. However, the skilled artisan will appreciate that in certain embodiments of the present invention, the sensor components are combined such that the plurality of components form one or more heterogeneous layers. In this context, those skilled in the art will understand that the order of the layered components may be varied in various embodiments of the present invention.

The embodiment shown in FIG. 1 includes a base layer 102 for supporting the sensor 100 . The base layer 102 may be made of materials such as metal and/or ceramic and/or polymeric substrates, which may be self-supporting or may be further supported by another material known in the art. Embodiments of the present invention include a conductive layer 104 disposed on and/or in combination with the base layer 102 . Typically, conductive layer 104 includes one or more electrodes. The operational sensor 100 typically includes a plurality of 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 that perform multiple functions, such as electrodes serving as both reference and counter electrodes.

As discussed in detail below, the base layer 102 and/or the conductive layer 104 may be created using a number of known techniques and materials. In certain embodiments of the invention, the circuitry of the sensor is defined by etching the provided conductive layer 104 into a desired pattern of conductive paths. A typical circuit for sensor 100 includes two or more adjacent conductive paths having regions at the proximal end to form contact pads and regions at the distal end to form sensor electrodes. An electrically insulating cover layer 106 such as a polymer coating may be provided over portions of the sensor 100 . Acceptable polymer coatings for use as insulating protective cover 106 may include, but are not limited to, non-toxic biocompatible polymers such as silicone compounds, polyimides, biocompatible solder masks, epoxy acrylates Copolymers, etc. In the sensor of the present invention, one or more exposed areas or apertures 108 may be formed through the cover layer 106 to open the conductive layer 104 to the external environment and, for example, to allow analytes such as glucose to penetrate the layers of the sensor and be sensed sensing element. The holes 108 may be formed by a variety of techniques, including laser ablation, tape masking, chemical grinding or etching or lithographic development, and the like. In certain embodiments of the present invention, a second photoresist may also be applied to the protective layer 106 during fabrication to define areas of the protective layer to be removed to form one or more holes 108 . The exposed electrodes and/or contact pads may also undergo secondary processing (eg, through holes 108 ) such as additional electroplating treatments to prepare surfaces and/or enhance conductive areas.

In the sensor configuration shown in FIG. 1, the analyte sensing layer 110 (which is typically a sensor chemical layer, meaning that the material in this layer undergoes a chemical reaction to produce a signal that can be sensed by the conductive layer) is disposed on the On one or more of the exposed electrodes of the conductive layer 104 are exposed. In the sensor configuration shown in Figure 2B, an interference suppression film 120 is disposed on one or more of the exposed electrodes of conductive layer 104, where the analyte sensing layer 110 is then disposed on the interference suppression film 120 on. 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 enzymes in the analyte sensing layer are combined with a second carrier protein such as human serum albumin, bovine serum albumin, and the like. In an illustrative embodiment, an oxidoreductase, such as glucose oxidase, in the analyte sensing layer 110 reacts with glucose to produce hydrogen peroxide, which compound then modulates the current at the electrodes. Since this modulation of the current depends on the concentration of hydrogen peroxide, which is related to the concentration of glucose, the glucose concentration can be determined by monitoring this modulation of the current. In particular embodiments of the present invention, hydrogen peroxide is oxidized at the working electrode (also referred to herein as the anode working electrode) as the anode, wherein the current produced is proportional to the hydrogen peroxide concentration. This modulation of electrical current caused by changes in hydrogen peroxide concentration can be monitored by any of a variety of sensor detector devices, such as general purpose sensor amperometric biosensor detectors or various others known in the art One of the similar devices, such as the glucose monitoring device produced by Medtronic MiniMed.

In embodiments of the invention, the analyte sensing layer 110 may be applied over a portion of the conductive layer 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 the counter electrode and/or the reference electrode. Although the thickness of the analyte sensing layer 110 can be as high as about 1000 micrometers (μm), the analyte sensing layer is typically relatively thin compared to the thicknesses found in sensors previously described in the art, and is typically less than 1 micrometer thick, for example , 0.5 micron, 0.25 micron or 0.1 micron. As discussed in detail below, some methods for producing a thin analyte sensing layer 110 include brushing the layer onto a substrate (eg, the reactive surface of a platinum black electrode), as well as spin coating processes, dip coating, and Drying process, low shear spraying process, inkjet printing process, silk screen printing process, etc.

Typically, the analyte sensing layer 110 is coated and/or disposed adjacent to one or more additional layers. Optionally, the 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, and the like. Typically, the protein layer 116 includes human serum albumin. In some embodiments of the invention, the additional layer includes an analyte modulating layer 112 disposed over the analyte sensing layer 110 to modulate the proximity of the analyte to the analyte sensing layer 110 . For example, the analyte modulating membrane layer 112 may include a glucose limiting membrane that regulates the amount of glucose in contact with an enzyme (eg, glucose oxidase) present in the analyte sensing layer. Such glucose-limiting membranes can be made from a variety of materials known to be suitable for such purposes, for example, silicone compounds such as polydimethylsiloxane, polyurethane, polyurea cellulose acetate, perfluorosulfonic acid ( NAFION), polyester sulfonic acid (eg Kodak AQ), hydrogels, polymer blends disclosed herein, or any other suitable hydrophilic membrane known to those skilled in the art.

In some embodiments of the present invention, as shown in FIG. 1, an adhesion promoter layer 114 is disposed between layers, such as the analyte modulating layer 112 and the analyte sensing layer 110, to facilitate their contact and/or or adhesion. In particular embodiments of the present invention, as shown in FIG. 1, an adhesion promoter layer 114 is disposed between the analyte modulating layer 112 and the protein layer 116 to facilitate their contact and/or adhesion. The adhesion promoter layer 114 may be made of any of a variety of materials known in the art to facilitate bonding between such layers. The adhesion promoter layer 114 typically includes a silane compound. In alternative embodiments, proteins or similar molecules in the analyte sensing layer 110 may be sufficiently cross-linked or otherwise prepared to allow the analyte modulating membrane layer 112 in the absence of the adhesion promoter layer 114 Disposed in direct contact with the analyte sensing layer 110 .

Examples of typical elements used to fabricate the sensors disclosed herein are discussed below.

Typical analyte sensor composition 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 (eg, layers), those skilled in the art will understand that sensors may be designed to contain elements/components (eg, elements that act as support substrate components and/or conductive components and/or components discussed below) or a combination of some or all of the material properties and/or functions for both the analyte-sensing component and the matrix for further use as an electrode in the sensor). It will be understood by those skilled in the art that these thin film analyte sensors can be adapted for use in many sensor systems, such as those described below.

base ingredient

Sensors of the present invention typically comprise a substrate component (see, eg, element 102 in Figure 1). The term "substrate component" is used herein according to art-recognized terminology and refers to a component in a device that typically provides a support matrix for a plurality of components stacked on top of each other and including functional sensors. In one form, the base composition comprises a thin film sheet of insulating (eg, electrically insulating and/or water impermeable) material. This substrate composition can be made of various materials with desired qualities such as dielectric properties, water impermeability, and air tightness. Some materials include metal and/or ceramic and/or polymer substrates and the like.

The base composition may be self-supporting or may be further supported by another material known in the art. In one embodiment of the sensor configuration shown in FIG. 1, the substrate composition 102 includes a ceramic. Alternatively, the base composition includes a polymeric material such as polyamide. In one illustrative embodiment, the ceramic substrate includes a composition that is primarily Al ₂ O ₃ (eg, 96%). The use of alumina as an insulating substrate component for use with implantable devices is disclosed in US Pat. Nos. 4,940,858, 4,678,868 and 6,472,122, which are incorporated herein by reference. The base composition of the present invention may further comprise other elements known in the art, such as sealing through holes (see eg WO 03/023388). Depending on the particular sensor design, the substrate composition may be a relatively thick composition (eg, thicker than 50, 100, 200, 300, 400, 500, or 1000 microns). Alternatively, non-conductive ceramics in thin compositions, such as alumina, can be utilized, eg, less than about 30 microns.

Conductive component

Electrochemical sensors of the present invention generally comprise a conductive component comprising at least one electrode for measuring the analyte to be determined or its by-products (eg oxygen and/or hydrogen peroxide) disposed on a substrate component ( See eg element 104 in Figure 1). The term "conductive component" is used herein according to art-recognized terminology and refers to a conductive sensor element, such as an electrode, capable of measuring a detectable signal and conducting the signal to a detection device. An illustrative example of the conductive component is a conductive component with a co-reactant that does not undergo an analyte (for use when the analyte interacts with a composition present in the analyte sensing component 110 (eg, the enzyme glucose oxidase) The conductive component may respond to exposure to a stimulus (eg, a change in concentration of an analyte or its by-products) compared to a reference electrode that changes in concentration of a reaction product of this interaction (eg, hydrogen peroxide) And measure the current increase or decrease. Illustrative examples of such elements include electrodes capable of producing variable detectable signals in the presence of variable concentrations of molecules such as hydrogen peroxide or oxygen. Typically, one of these electrodes in the conductive composition is the working electrode, which may be made of a non-corrosive metal or carbon. Carbon working electrodes can be vitreous or graphitic and can be made from solids or pastes. Metal working electrodes can be made of platinum group metals, including palladium or gold, or non-corrosive metal conductive oxides, such as ruthenium dioxide. Alternatively, the electrode may comprise a silver/silver chloride electrode composition. The working electrode may be a wire or a thin conductive film applied to the substrate, for example by coating or printing. Typically, only a portion of the metal or carbon conductor surface is in electrolytic contact with the analyte-containing solution. This part is called the working surface of the electrode. The remaining surfaces of the electrodes are typically isolated from the solution by an electrically insulating cover composition 106 . Examples of useful materials for creating this protective covering composition 106 include polymers such as polyimide, polytetrafluoroethylene, polyhexafluoropropylene, and silicones such as polysiloxane.

In addition to the working electrode, the analyte sensors of the present invention typically comprise a reference electrode or a combined reference and 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 contain a separate counter electrode, which may be made of the same or a different material than the working electrode. Typical sensors of the present invention have 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 or four or more working electrodes. The working electrodes in the sensor can be connected integrally or they can also be kept separate.

Typically, for in vivo use, embodiments of the present invention are implanted subcutaneously into the skin of a mammal for direct contact with the mammal's bodily fluids, such as blood. Alternatively, the sensor can be implanted in other areas of the mammal's body, such as the intraperitoneal space. When multiple working electrodes are used, the multiple working electrodes may be implanted together or at different locations in the body. The counter electrode, reference electrode and/or counter/reference electrode may also be implanted near one or more working electrodes or at other locations within the mammalian body. Embodiments of the present invention include sensors that include electrodes composed of nanostructured materials. As used herein, a "nanostructured material" is an object fabricated to have at least one dimension 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, and the like.

interference suppressor

The electrochemical sensor of the present invention optionally contains an interference suppressing component disposed between the electrode surface and the environment to be measured. In particular, certain sensor embodiments rely on the oxidation and/or reduction of hydrogen peroxide produced by an enzymatic reaction on the surface of the working electrode at a constant applied potential. Because amperometric detection based on direct oxidation of hydrogen peroxide requires a relatively high oxidation potential, sensors employing this detection scheme may experience interference from oxidizable species such as ascorbic acid, uric acid, and acetamidophenol present in biological fluids. In this context, the term "interference suppressing component" is used herein according to art-recognized terminology and refers to a coating or film in a sensor that acts to suppress spurious signals generated by such oxidizable species, The spurious signal interferes with the detection of the signal produced by the analyte to be sensed. Certain interference suppressing components act by size exclusion (eg, by excluding interfering species of a particular size). Examples of interference suppressing ingredients include one or more layers or coatings of compounds such as hydrophilic cross-linked pHEMA and lysine polymers and cellulose acetate (including cellulose acetate incorporating agents such as poly(ethylene glycol) , polyethersulfone, polytetrafluoroethylene, perfluoro ionomer NAFION, polyphenylenediamine, epoxy resin, etc.). Illustrative discussions of such interference suppressing components exist, 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, interference-suppressing components that are observed to move based on molecular weight range limiting compounds, such as cellulose acetate as disclosed, for example, in US Pat. No. 5,755,939, the contents of which are incorporated by reference. Disclosed herein, for example, in US Patent Application Serial No. 12/572,087, are such compositions and other compositions having a range of undesired material properties that may make additional compositions ideal for use as interference suppression membranes in certain amperometric glucose sensors Method of manufacture and use.

Analyte Sensing Components

Electrochemical sensors of the present invention comprise analyte-sensing components (see, eg, element 110 in FIG. 1 ) disposed on electrodes of the sensor. The term "analyte sensing component" is used herein in accordance with art-recognized terminology and is meant to include a component capable of identifying or reacting with an analyte whose presence is to be detected by an analyte sensor device. Typically, this material in the analyte-sensing component, upon interaction with the analyte to be sensed, produces a detectable signal, typically via the electrodes of the conductive component. In this regard, the electrodes of the analyte sensing component and the conductive component operate in combination to generate electrical signals that are read by a device associated with the analyte sensor. Typically, analyte-sensing components include oxidoreductases (eg, glucose oxidase) capable of reacting with and/or causing concentration changes of the molecules, as can be achieved by measuring conductive components (eg, oxygen and/or hydrogen peroxide) The current change at the electrode is used to measure the concentration change of the molecule. Enzymes capable of producing molecules such as hydrogen peroxide can be placed on electrodes according to a number of methods known in the art. The analyte-sensing composition may coat all or a portion of each electrode of the sensor. In this context, the analyte-sensing composition may coat the electrodes to an equivalent degree. Alternatively, the analyte-sensing components may coat different electrodes to varying degrees, with, for example, the coated surface of the working electrode being larger than the coated surface of the counter and/or reference electrode.

A typical sensor embodiment of this element of the invention utilizes an enzyme (eg glucose oxidase) that has been combined in a fixed ratio with a second protein (eg albumin) and then applied to the surface of the electrode to form a thin enzymatic component (eg typically for An enzyme optimized for glucose oxidase stabilization properties). In a typical embodiment, the analyte sensing component includes a GOx and HSA mixture. In a typical embodiment of an analyte-sensing component with GOx, GOx reacts with glucose present in the sensing environment (eg, the body of a mammal) and produces hydrogen peroxide according to the reaction shown in FIG. 1 , wherein The hydrogen peroxide produced is detected at the anode at the working electrode in the conductive component.

As mentioned above, the enzyme and second protein (eg, albumin) are typically treated to form a cross-linked matrix (eg, by adding a cross-linking agent to the protein mixture). As is known in the art, cross-linking conditions can be manipulated to modulate factors such as the retained biological activity of the enzyme, its mechanical and/or operational stability. Illustrative cross-linking procedures are described in US Patent Application Serial No. 10/335,506 and PCT Publication WO 03/035891, which are incorporated herein by reference. For example, amine cross-linking reagents (such as, but not limited to, glutaraldehyde) can be added to the protein mixture.

protein composition

The electrochemical sensor of the present invention optionally includes a protein component (see eg, element 116 in Figure 1) disposed between the analyte sensing component and the analyte modulating component. The term "protein component" is used herein according to art-recognized terminology, and refers to a component comprising carrier proteins, etc., selected to be compatible with the analyte sensing component and/or the analyte modulating component. In typical embodiments, the protein component includes albumin, such as human serum albumin. The HSA concentration can vary between about 0.5%-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 to replace or supplement HSA. This component is typically cross-linked on the analyte sensing component according to art-recognized protocols.

adhesion promoting ingredients

Electrochemical sensors of the present invention may contain one or more adhesion promoting (AP) components (see, eg, element 114 in Figure 1). The term "adhesion promoting component" is used herein in accordance with art-recognized terminology and refers to a component comprising a material selected to promote adhesion between adjacent components in a sensor. The adhesion promoting component is typically disposed between the analyte sensing component and the analyte modulating component. The adhesion promoting component is typically disposed between the optional protein component and the analyte modulating component. The adhesion promoter component can be made from any of a variety of materials known in the art for promoting bonding between such components, and can be made by any of a variety of methods known in the art method to apply. The adhesion promoter component typically includes a silane compound such as gamma-aminopropyltrimethoxysilane.

The use of silane coupling agents, particularly those of the formula R'Si(OR) ₃ , where R' is typically an aliphatic group with a terminal amine and R is a lower alkyl group, to promote adhesion is known in the art. (see, eg, US Pat. No. 5,212,050, which is incorporated herein by reference). For example, where silanes such as gamma-aminopropyltriethoxysilane and glutaraldehyde are used in a stepwise method to allow bovine serum albumin (BSA) and glucose oxidase ( _GOx ) to adhere to the electrode surface and interact with the electrode Chemically modified electrodes with surface co-crosslinking are known in the art (see eg Yao, T. Acta Analytical Chemistry 1983, 148, 27-33).

In certain embodiments of the invention, the adhesion promoting component further comprises one or more compounds that may also be present in adjacent components for limiting diffusion of an analyte, such as glucose, through the analyte modulating component, whereby said adjacent components such as polydimethylsiloxane (PDMS) compounds. In illustrative embodiments, the formulation includes 0.5-20% PDMS, typically 5-15% PDMS, and most typically 10% PDMS. In certain embodiments of the invention, the adhesion-promoting component is cross-linked within the layered sensor system and correspondingly comprises selected for its ability to cross-link moieties present in proximal components such as analyte-modulating components Pharmacy. In illustrative embodiments of the invention, the adhesion promoting component comprises an agent for which the amine or carboxyl moiety of a protein present in a proximal component such as an analyte sensing component and/or a protein component is present The ability to crosslink and/or crosslink siloxane moieties present in compounds disposed in proximal layers such as the analyte modulating layer is selected.

Analyte Modulator

Electrochemical sensors of the present invention comprise an analyte modulating component (see, eg, element 112 in Figure 1) disposed on the sensor. The term "analyte modulating component" is used herein according to art-recognized terminology and refers to a component that typically forms a membrane on a sensor that operates to modulate the passage of one or more analytes (eg, glucose) through Diffusion of the ingredients. In certain embodiments of the invention, the analyte-modulating component is an analyte-limiting membrane (eg, a glucose-limiting membrane) that operates to prevent or restrict passage of one or more analytes (eg, glucose) through Diffusion of the ingredients. In other embodiments of the invention, the analyte modulation component operates to facilitate diffusion of one or more analytes through the component. Optionally, such analyte-modulating components can be formed to prevent or restrict diffusion of one type of molecule (eg, glucose) through the component, while allowing or even facilitating the diffusion of other types of molecules (eg, O ₂ ) therethrough the ingredients. Typically, the analyte modulating component includes a polycarbonate polymer composition as disclosed herein.

With regard to glucose sensors, in known enzyme electrodes, glucose and oxygen from the blood and some interfering substances such as ascorbic acid and uric acid diffuse across the primary membrane of the sensor. When glucose, oxygen, and interferents reach the analyte detection component, enzymes such as glucose oxidase catalyze the conversion of glucose to hydrogen peroxide and gluconolactone. The hydrogen peroxide can diffuse back through the analyte modulating component, or it can diffuse to an 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 therethrough. In this case, the illustrative analyte modulating component is a semi-permeable membrane that allows the passage of water, oxygen, and at least one selective analyte and has the ability to absorb water, the membrane having a water-soluble hydrophilic Sexual polymers.

Various illustrative analyte-modulating compositions are known in the art and described, for example, in US Pat. Nos. 6,319,540, 5,882,494, 5,786,439, 5,777,060, 5,771,868, and 5,391,250, each of which The disclosure of 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 an exemplary embodiment of the present invention, the analyte modulating composition comprises the polycarbonate polymeric composition disclosed herein.

cover ingredients

Electrochemical sensors of the present invention comprise one or more covering components (see, eg, element 106 in FIG. 1 ), which are typically electrically insulating protective components. Typically, such a covering component may be in the form of a coating, sheath or tube and disposed over at least a portion of the analyte modulating component. Acceptable polymer coatings for use as insulating protective covering components may include, but are not limited to, non-toxic biocompatible polymers such as silicone compounds, polyimides, biocompatible solder masks, epoxy acrylics Ester copolymers, etc. Further, these coatings may be photoimageable to facilitate lithographic formation of holes through the conductive components. A typical covering composition includes spinning on silicone. This ingredient may be a commercially available RTV (room temperature vulcanizing) silicone composition, as known in the art. A typical chemical in this context is polydimethylsiloxane (based on acetoxy groups).

Illustrative Embodiments of Analyte Sensor Devices and Associated Properties

The analyte sensor devices disclosed herein have many embodiments. A general embodiment of the present invention is an analyte sensor device for implantation in a mammal. 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 , joint fluid, gastric juice, etc. Additionally, solid or dry samples can be dissolved in a suitable solvent to provide a liquid mixture suitable for analysis.

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 certain embodiments, as typically occurs with subcutaneous sensors, the sensor may be in direct contact with the interstitial fluid. The sensors of the present invention may 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, eg, US Pat. Nos. 6,155,992 and 6,706,159, which are incorporated herein by reference). In other embodiments, the sensor may be in contact with blood, as typically occurs with intravenous sensors, for example. Sensor embodiments of the present invention further include sensors suitable for use in various situations. For example, in some embodiments, the sensor may be designed for use in mobile situations, as employed by mobile users. Alternatively, the sensor may be designed for use in a stationary situation, such as is suitable for use in a clinical setting. Such sensor embodiments include, for example, sensors for monitoring one or more analytes present in one or more physiological environments in hospitalized patients.

The sensors of the present invention may also be incorporated into a wide variety of medical systems known in the art. The sensors of the present invention can be used, for example, in closed-loop infusion systems designed to control the rate of drug infusion into the body of a user. Such closed-loop infusion systems may contain sensors and associated meters that generate inputs to a controller, which in turn operates a delivery system (eg, a system that calculates a dose to be delivered by a drug infusion pump). In such cases, meters associated with the sensors may also transmit commands to the delivery system and may be used to remotely control the delivery system. Typically, the sensor is a subcutaneous sensor that is in contact with the interstitial fluid to monitor the glucose concentration in the user, and the fluid infused into the user by the delivery system contains insulin. Illustrative systems are disclosed, for example, in US Patent Nos. 6,558,351 and 6,551,276; PCT Application Nos. US99/21703 and US99/22993; and WO 2004/008956 and WO 2004/009161, all by reference Incorporated herein.

Certain embodiments of the present invention measure peroxide and have advantageous properties suitable for implantation in a variety of sites in mammals, including areas for subcutaneous and intravenous implantation and implantation into various non- in the vascular area. Peroxide sensor designs that allow implantation in non-vascular regions have advantages over certain sensor device designs that measure oxygen due to oxygen noise issues that can arise in oxygen sensors implanted in non-vascular regions. For example, in such implantable oxygen sensor device designs, oxygen noise at the reference sensor can compromise the signal-to-noise ratio, interfering with the implantable oxygen sensor device's ability to obtain stable glucose readings in this environment. Thus, the peroxide sensor of the present invention overcomes the difficulties observed with such oxygen sensors in non-vascular regions.

Certain peroxide sensor embodiments of the present invention further comprise advantageous long-term or "permanent" sensors suitable for implantation in mammals for periods of time greater than 30 days. Specifically, as known in the art (see, eg, ISO 10993, "Biological Evaluation of Medical Devices"), medical devices of sensors as described herein can be divided into three groups based on the duration of implantation : (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 according to this classification, ie >30 days. In related embodiments of the present invention, the highly stable design of the peroxide sensor of the present invention allows the implantable sensor to continue to operate in this regard for 2 months, 3 months, 4 months, 5 months, 6 months or 12 or more months.

Arrangement of Analyte Sensor Devices and Elements

As mentioned above, the invention disclosed herein encompasses a number of embodiments, including a sensor having a series of elements comprising a polycarbonate polymeric film. Such embodiments of the present invention allow the skilled artisan to create various arrangements of the analyte sensor devices disclosed herein. As noted above, illustrative general embodiments of the sensors disclosed herein include a base layer, a cover layer, and at least one layer having sensor elements, such as electrodes, disposed between the base layer and the cover layer. Typically, exposed portions of one or more sensor elements (eg, working electrode, counter electrode, reference electrode, etc.) are coated with a very thin layer of material with the appropriate electrode chemistry. For example, enzymes such as lactate oxidase, glucose oxidase, glucose dehydrogenase, or hexokinase can be disposed on exposed portions of the sensor element within openings or pores defined in the cover layer. Figure 1 shows a cross-section of a typical sensor structure 100 of the present invention. According to the method of the present invention for producing the sensor structure 100, the sensor is formed from multiple layers of various conductive and non-conductive components disposed on each other.

As described above, in the sensors of the present invention, various layers of the sensor (eg, analyte sensing layers) may have one or more bioactive and/or inert materials incorporated therein. The term "incorporated" as used herein is meant to describe any state or condition in which the incorporated material is held on the outer surface of the layer or within the solid phase or support matrix of the layer. Thus, "incorporated" materials may, for example, be immobilized, physically entrapped, covalently adhered to functional groups of one or more matrix layers. Furthermore, these additional steps or agents may be employed if any method, agent, additive or molecular linker that facilitates the "incorporation" of the material is not detrimental to the invention but is consistent with the purpose of the invention. Of course, this definition applies to any of the embodiments of the invention in which a biologically active molecule (eg, an enzyme such as glucose oxidase) is "incorporated". For example, certain layers of the sensors disclosed herein contain proteinaceous species, such as albumin, that serve as a crosslinkable matrix. As used herein, proteinaceous material is meant to encompass materials generally derived from proteins, whether the actual material is native protein, inactivated protein, denatured protein, hydrolyzed material, or a derivative thereof. Examples of suitable proteinaceous materials include, but are not limited to, enzymes such as glucose oxidase and lactate oxidase, albumin (eg, human serum albumin, bovine serum albumin, etc.), casein, gamma-globulin, collagen, and collagen derivatives (eg, fish gelatin, isinglass, animal gelatin, and animal glue).

An illustrative embodiment of the present invention is shown in FIG. 1 . This embodiment includes an electrically insulating base layer 102 for supporting the sensor 100 . The electrically insulating base layer 102 may be made of a material such as a ceramic substrate, which 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 can facilitate clean, high-density mass production. Further, in some production processes using such polyimide tapes, the sensor 100 may be produced on both sides of the tape.

Exemplary embodiments of the present invention include an analyte sensing layer disposed on substrate layer 102 . In the illustrative embodiment shown in FIG. 1 , the analyte sensing layer includes a conductive layer 104 disposed on an insulating base layer 102 . Typically, conductive layer 104 includes one or more electrodes. As described below, the conductive layer 104 may 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 pattern of conductive paths. A typical circuit for sensor 100 includes two or more adjacent conductive paths having regions at the proximal end to form contact pads and regions at the distal end to form sensor electrodes. An electrically insulating protective cover layer 106 , such as a polymer coating, may typically be disposed over portions of the conductive layer 104 . Acceptable polymer coatings for use as insulating protective layer 106 may include, but are not limited to, non-toxic biocompatible polymers such as polyimides, biocompatible solder masks, epoxy acrylate copolymers, and the like. Further, these coatings may be photoimageable to facilitate lithographic formation of holes 108 through conductive layer 104 . In certain embodiments of the present invention, the analyte sensing layer is disposed on a porous metal and/or ceramic and/or polymer matrix, wherein this combination of elements is used as an electrode in a sensor.

In the sensor of the present invention, one or more exposed areas or holes 108 may be made through the protective layer 106 to the conductive layer 104 to define the contact pads and electrodes of the sensor 100 . In addition to photolithographic development, apertures 108 can be formed by a number of techniques, including laser ablation, chemical grinding, or etching, among others. A second photoresist may also be applied to the capping layer 106 to define areas where the protective layer is to be removed to form the holes 108 . Operational sensor 100 typically includes a plurality of electrodes, such as a working electrode and a counter electrode, that are electrically insulated from each other, yet typically positioned in close proximity 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 treatments, to prepare surfaces and/or enhance conductive areas.

The analyte sensing layer 110 is generally disposed on one or more of the exposed electrodes of the conductive layer 104 through the apertures 108 . The analyte sensing layer 110 is typically a sensor chemical layer, and most typically 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 produce hydrogen peroxide that modulates a current to the electrodes, which can be monitored to measure the amount of glucose present. The sensor chemistry layer 110 may 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 the portions of the working and counter electrodes that include the conductive layers. Some methods for creating the thin sensor chemistry layer 110 include spin coating processes, soak and dry processes, low shear spray coating processes, ink jet printing processes, screen processes, and the like. Most typically, the thin sensor chemistry layer 110 is applied using a spin coating process.

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 can regulate the amount of analyte that can contact the enzyme of the analyte sensing layer. For example, the coating may include an analyte modulating membrane layer, such as a glucose limiting membrane, that regulates the amount of glucose on the electrode that contacts the glucose oxidase layer. Such glucose limiting membranes can be made from a wide variety of materials known to be suitable for such purposes, such as silicone, polyurethane, polyurea cellulose acetate, perfluorosulfonic acid, polyestersulfonic acid ( Kodak AQ), hydrogels or any other membrane known to those skilled in the art. In certain embodiments of the present invention, the analyte modulating layer comprises a linear polyurethane/polyurea polymer polycarbonate with a branched acrylate hydrophilic comb copolymer, the branched acrylate hydrophilic comb The copolymer has a central chain and a plurality of side chains coupled to the central chain, wherein at least one of the side chains includes a silicone moiety.

In some embodiments of the invention, the coating is a glucose limiting membrane layer 112 disposed over the sensor chemical layer 110 to regulate the contact of glucose with the sensor chemical layer 110 . In some embodiments of the present invention, as shown in FIG. 1, an adhesion promoter layer 114 is disposed between the membrane layer 112 and the sensor chemical layer 110 to facilitate their contact and/or adhesion. The adhesion promoter layer 114 may be made of any of a variety of materials known in the art to facilitate bonding between such layers. The adhesion promoter layer 114 typically includes a silane compound. In alternative embodiments, the proteins or similar molecules in sensor chemistry layer 110 may be sufficiently cross-linked or otherwise prepared to allow membrane layer 112 to be placed in contact with sensor chemistry in the absence of adhesion promoter layer 114 Layer 110 is in direct contact.

As described above, embodiments of the present invention may include one or more functional coatings. The term "functional coating" as used herein means coating at least a portion of at least one surface of a sensor, and more generally coating substantially all of the surface of the sensor and capable of interacting with one or more of the sensors in the environment in which the sensor is disposed A layer at which analytes interact, such as one or more analytes such as compounds, cells and fragments thereof, and the like. Non-limiting examples of functional coatings include sensor chemical layers (eg, enzyme layers), analyte confinement layers, biocompatible layers; layers that increase the slipperiness of the sensor; layers that promote connection to sensor cells; connected layers; etc. Typically, the analyte modulating layer operates to prevent or limit the diffusion of one or more analytes, such as glucose, through the layer. Optionally, such a layer can be formed to prevent or restrict the diffusion of one type of molecule (eg glucose) through the layer, while allowing or even facilitating the diffusion of other types of molecules (eg _O2 ) through the layer . Illustrative functional coatings are hydrogels, such as those disclosed in US Pat. Nos. 5,786,439 and 5,391,250, the disclosures of which are 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 layer.

The sensor embodiments disclosed herein may include layers having UV absorbing polymers. According to one aspect of the present invention, there is provided a sensor comprising at least one functional coating comprising a UV absorbing polymer. In some embodiments, the UV absorbing polymer is a polyurethane, polyurea or polyurethane/polyurea copolymer. More typically, the UV absorbing polymer of choice 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.

UV absorbing polymers are advantageously used in various sensor fabrication methods, such as those described in: US Patent No. 5,390,671 to Lord et al, entitled "Transcutaneous Sensor Insertion Set" ; No. 5,165,407, entitled "Implantable Glucose Sensor," to Wilson et al.; and "Two-Dimensional Diffusion Glucose Substrate Sensing Electrode," to Gough et al. Sensing Electrode)" US Patent No. 4,890,620, which is incorporated herein by reference in its entirety. However, any method of sensor production that includes the step of forming a UV absorbing polymer layer over or under the sensor element is considered within the scope of the present invention. In particular, the method of the present invention is not limited to thin film fabrication methods and can work with other sensor fabrication methods utilizing UV laser cutting. Embodiments can work with thick film, planar or cylindrical sensors, etc., as well as other sensor shapes that require laser cutting.

As disclosed herein, the sensors of the present invention are specifically designed to be used as subcutaneous or transcutaneous glucose sensors to monitor blood glucose levels in diabetic patients. Typically, each sensor includes a plurality of sensor elements, eg, conductive elements, such as elongated thin film conductors, formed between an underlying insulating film base layer and an overlying insulating film cover layer.

If desired, multiple different sensor elements may be included in a single sensor. For example, both conductive sensor elements 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 a plurality of openings or holes in its cover layer. One or more openings may also be defined in the cover layer directly over a portion of the base layer to provide for interaction of the base layer with one or more analytes in the environment in which the sensor is disposed. The base layer and cover layer can comprise a variety of materials, usually polymers. In a more specific embodiment, the base layer and the cover layer comprise insulating materials such as polyimide. Openings are typically formed in the cover layer to expose the distal electrodes and proximal contact pads. For example, in glucose monitoring applications, the sensor may be placed percutaneously such that the distal electrode is in contact with the patient's blood or extracellular fluid, and the contact pads are placed externally for easy connection to the monitoring device.

Analyte Sensor Device Configuration

In a clinical setting, electrochemical sensors can be used to determine accurate and relatively rapid measurements such as glucose and/or lactate levels from blood samples. Conventional sensors are manufactured as large, including many serviceable components, or as small, planar sensors that may be more convenient in many situations. The term "planar" as used herein refers to well-known procedures for fabricating substantially planar structures comprising relatively thin layers of material, eg, using well-known thick film or thin film techniques. See, eg, US Patent No. 4,571,292 to Liu et al. and US 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 (eg, planar) than existing sensors in the prior art. Additionally, certain embodiments of the present invention include one or more of the sensors disclosed herein coupled to another device, such as a drug infusion pump.

Figure 2 provides a diagrammatic view of a typical analyte sensor configuration of the present invention. Certain sensor configurations are in a relatively flat "ribbon" type configuration that can be fabricated with analyte sensor devices. This "ribbon" type configuration demonstrates the advantages of the sensors disclosed herein due to spin-coating of sensing enzymes such as glucose oxidase, which results in highly flexible sensors that allow design and production Manufacturing steps for extremely thin enzymatic coatings of geometric shapes. Such thin enzyme-coated sensors provide additional advantages, such as allowing a smaller sensor area while maintaining sensor sensitivity, a highly desirable feature of implantable devices (eg, smaller devices are easier to implant). Thus, sensor embodiments of the present invention that utilize very thin analyte sensing layers, which may be formed by processes such as spin coating, may have a wider range of Geometric configuration (eg plane).

Certain sensor configurations contain multiple conductive elements, such as multiple working, counter, and reference electrodes. The advantages of such a configuration include increased surface area, thereby providing higher sensor sensitivity. For example, one sensor configuration introduces a third working sensor. A clear advantage of this type of configuration is that the signals of the three sensors are averaged, thereby increasing the accuracy of the sensors. Other advantages include the ability to measure multiple analytes. Specifically, an analyte sensor configuration that includes electrodes in this arrangement (eg, multiple working, counter, and reference electrodes) can be incorporated 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 many advantages such as the linear response and ease of calibration and/or Recalibration.

An exemplary multi-sensor device includes a single device having a first sensor that is cathodically polarized and designed to measure glucose at the working electrode (cathode) due to the interaction of glucose with glucose oxidase The change in oxygen concentration that occurs; and a second sensor that is anodically polarized and designed to measure the process that occurs at the working electrode (anode) due to glucose from the external environment and interaction with glucose oxidase; Changes in hydrogen oxide concentration. As is known in the art, in such designs, the first oxygen sensor typically experiences a reduction in current at the working electrode when oxygen touches the sensor, and when the hydrogen peroxide produced as shown in Figure 1 touches the sensor, The second hydrogen peroxide sensor typically experiences an increase in current at the working electrode. Additionally, as is known in the art, observations of current changes occurring at the working electrode are correlated with changes in the concentrations of oxygen and hydrogen peroxide molecules, which are then compared to the reference electrode in the corresponding sensor system. Changes in the concentration of hydrogen oxide molecules may be related to the glucose concentration in the external environment (eg, mammalian bodies).

The analyte sensors of the present invention can be coupled to other medical devices such as drug infusion pumps. In an illustrative variation of this approach, the replaceable analyte sensor of the present invention may be coupled to other medical devices such as drug infusion pumps, for example, through a port (eg, a subcutaneous port with a locking electrical connection) coupled to the medical device.

Illustrative methods and materials for making the analyte sensor devices of the present invention

Numerous articles, US patents, and patent applications describe the prior art with the common methods and materials disclosed herein, and further describe various elements (and methods of making them) that can be used in the sensor designs disclosed herein. These include, for example, US Patent Nos. 6,413,393; 6,368,274; 5,786,439; 5,777,060; 5,391,250; 5,390,671; US Patent Application No. 20020090738; and PCT International Publication Nos. WO 01/58348, 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 and WO 03/074107, said documents The contents of each document in are incorporated herein by reference.

A typical sensor for monitoring glucose concentration in diabetic patients is further described in "In VivoCharacteristics of Needle-Type Glucose Sensor-Measurements of Subcutaneous Glucose Concentrations in Human Volunteers" by Shichiri et al. in Human Volunteers)”, Horm. Metab. Res., Suppl. 20:17-20 (1988); Bruckel et al.: “Using an enzymatic glucose sensor and the Vik method on subcutaneous glucose concentrations 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., "Diabetes In Vivo Molecular Sensing in DiabetesMellitus: An Implantable Glucose Sensor with Direct Electron Transfer", Diabetologia 32:213-217 (1989). Other sensors are described, for example, in Reach et al., ADVANCES IN IMPLANTABLE DEVICES, A. Turner (ed.), JAI Press, London, Chapter 1, (1993), which is incorporated by reference. into this article.

An exemplary embodiment of the invention disclosed herein is a method of fabricating a sensor device for implantation in a mammal, the method comprising the steps of: providing a substrate layer; forming a conductive layer on the substrate layer, wherein the conductive layer includes electrodes (and typically working, reference, and counter electrodes); forming an analyte-sensing layer on a conductive layer, wherein the analyte-sensing layer comprises a composition that alters conductivity in the presence of the analyte current at electrodes in layers; optionally forming a protein layer on the analyte sensing layer; forming an adhesion promoting layer on the analyte sensing layer or optional protein layer; forming an assay disposed on the adhesion promoting layer an analyte modulating layer, wherein the analyte modulating layer comprises a composition that modulates the diffusion of the analyte therethrough; and forming a capping layer disposed over at least a portion of the analyte modulating layer, wherein the capping layer further comprises an analyte modulating layer A hole over at least a portion of the layer. In certain embodiments of the present invention, the analyte modulating layer comprises a linear polyurethane/polyurea polymer polycarbonate with a branched acrylate copolymer having a central chain and coupled to the central Multiple sidechains of the chain. In some embodiments of these methods, the analyte sensor device is formed in a planar geometric configuration.

As disclosed herein, the various layers of the sensor can be fabricated to exhibit a variety of different properties that can be manipulated according to the particular design of the sensor. For example, the adhesion promoting layer contains a compound selected for its ability to stabilize the entire sensor structure, typically a silane composition. In some embodiments of the invention, the analyte sensing layer is formed by a spin coating process and has a thickness selected from the group consisting of less than 1 micron, 0.5 microns, 0.25 microns, and 0.1 microns in height.

The method of making a sensor generally includes the step of forming a protein layer on the analyte sensing layer, wherein the protein within the protein layer is an albumin selected from the group consisting of bovine serum albumin and human serum albumin. The method of making a sensor generally includes 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 includes a carrier protein composition in 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.

The electrodes of the present invention can be formed from a wide variety of materials known in the art. For example, the electrodes may be made of noble late transition metals. Metals such as gold, platinum, silver, rhodium, iridium, ruthenium, palladium or osmium may be suitable for various embodiments of the present invention. Other compositions such as carbon or mercury are also useful in certain sensor embodiments. Among these metals, silver, gold or platinum is usually used as the reference electrode metal. A subsequently chlorinated silver electrode is usually used as a reference electrode. These metals can be deposited by any means known in the art, including the plasma deposition methods cited above, or by electroless methods, which may involve when the substrate is immersed in a solution containing a metal salt and a reducing agent metal is deposited onto previously metallized areas. The electroless process continues as the reducing agent donates electrons to the conductive (metallized) surface, with concomitant reduction of the metal salt at the conductive surface. The result is an adsorbed metal layer. (For an additional discussion of electroless methods, see: Wise, E.M. 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. "None Electroless Plating," Modern Electroplating, Lowenheim, F.A., ed., Wiley, New York, N.Y. (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 to the catalytic metal electrode surface. Such high density of active sites is a necessary property for efficient redox conversion of electroactive species such as hydrogen peroxide.

In an exemplary embodiment of the present invention, the base 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 provided as a plurality of thin film conductive layers, such as an initial chromium-based layer suitable for chemical adhesion to a polyimide base layer, followed by a sequentially formed gold-based thin film layer and chromium-based thin film layers. 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 photolithographic techniques, and a contact mask can be applied over the photoresist coating for suitable photoimaging. The contact mask typically contains one or more conductor trace patterns for properly exposing the photoresist coating, followed by an etching step to leave the plurality of conductive sensor traces on the base 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 (eg, working, counter, and reference electrodes).

Portions of the sensor conductive layer are typically covered by an insulating cover layer, typically a material such as silicon polymer and/or polyimide. The insulating cover layer can be applied in any desired manner. In an exemplary procedure, an insulating cover layer is applied as a liquid layer over the sensor traces, after which the substrate is rotated to distribute the liquid material as a thin film over the sensor traces and extend the liquid material as a thin film over the sensor traces beyond the edge margin of hermetic contact with the base layer. This liquid material may 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 may be applied using spray techniques or any other desired means of application. Various insulating layer materials can be used, such as photoimageable epoxy acrylate, with an illustrative material including photoimageable polyimide available from OCG, West Paterson, NJ, Product No. 7020.

Optionally, as described above, after exposing the sensor tip through the opening, an appropriate electrode chemistry that defines the distal electrode can be applied to the sensor tip. In an illustrative sensor embodiment with three electrodes for use as a glucose sensor, an enzyme (usually glucose oxidase) is disposed within one of the openings, thus coating one of the sensor tips to define the working electrode. One or both of the other electrodes can be provided with the same coating as the working electrode. Alternatively, other suitable chemicals such as other enzymes may be provided uncoated to the other two electrodes, or provided with chemicals to define the reference and counter electrodes of the electrochemical sensor.

Methods for producing the extremely thin enzymatic coatings of the present invention include spin coating processes, soaking and drying processes, low shear spraying processes, ink jet printing processes, screen processes, and the like. Because a skilled artisan can readily determine the thickness of an enzymatic coating applied by processes in the art, a skilled artisan can readily identify methods capable of producing the extremely thin coatings of the present invention. Typically, such coatings are steam crosslinked after their 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 improved signal-to-noise ratio. Additionally, embodiments of the invention utilizing glucose oxidase coatings formed by such processes are designed to recycle hydrogen peroxide and improve the biocompatibility profile of such sensors.

Sensors produced by processes such as spin coating also avoid other problems associated with electrodeposition, such as those associated with material stress placed on the sensor during the electrodeposition process. In particular, the electrodeposition process was observed to generate mechanical stress on the sensor, eg, from tensile 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 was not observed in coatings placed on the sensor by spin coating or other low stress processes. Accordingly, yet another embodiment of the present invention is a method of avoiding electrodeposition-affected cracking and/or delamination of a coating on a sensor, the method comprising applying the coating by a spin coating process.

Methods for using the analyte sensor device of the present invention

A related embodiment of the present invention is a method of sensing an analyte in a mammal, the method comprising implanting the analyte sensor embodiments disclosed herein into the mammal, and then sensing a change in current at a working electrode As well as correlating the current change with the presence of the analyte, allowing the analyte to be sensed. The analyte sensor can be anodically polarized such that the working electrode that senses the change in current is the anode, or the working electrode that senses the change in current is the cathode. In one such method, the analyte sensor device senses mammalian glucose. In alternative embodiments, the analyte sensor device senses lactate, potassium, calcium, oxygen, pH, and/or any physiologically relevant analyte in the mammal.

Certain analyte sensors having the structures discussed above have many highly desirable properties 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 an analyte in the mammal for more than 1 month, more than 2 months, more than 3 months, more than 4 months, More than 5 months or more than 6 months. Typically, an analyte sensor device so implanted in a mammal senses a change in current in response to the analyte within 15, 10, 5, or 2 minutes of the analyte contacting the sensor. In such methods, sensors can be implanted in various locations within the mammalian body, such as in both vascular and non-vascular spaces.

example

The following examples are given to assist in understanding the invention, but it is to be understood that this invention is not limited to the specific materials or procedures of the examples. All materials used in the examples were obtained from commercial sources.

Example 1: Synthesis and Characterization of Illustrative Polyurea/Polyurethane Polymers Using Conventional Methods:

The disclosure provided herein, combined with what is known in the art, demonstrates that functional linear polyurethane/polyurea polymers can be made from a number of formulations, such as those disclosed in: US Pat. Nos. 5,777,060; 5,882,494; 6,642,015; and PCT Publication Nos. WO 96/30431; WO 96/18115; WO 98/13685; and WO 98/17995, the contents of which are incorporated herein by reference. Certain of these polymers provide formulations useful as glucose limiting membranes (GLMs).

Standard GLM formulations used to make 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 having a molecular weight of approximately 600); and 50 mol% diisocyanate (eg, 4,4'-diisocyanate).

Such standard GLM formulations and methods for their synthesis are disclosed, for example, in US Pat. Nos. 6,642,015, 5,777,060 and 6,642,015.

Another formulation used in the examples of the present invention is referred to as "semi-osmotic GLM" because its glucose permeability was observed to be half that of the standard formulation described above. In standard GLM, Jeffamine/PDMS quantification = 3/1 (molar ratio). In contrast, in a "semi-permeable GLM" this ratio is changed such that Jeffamine/PDMS=12/1. Such semi-permeable GLMs can be used, for example, to reduce the weight % of GLM urea in the entire polymer blend to achieve a specific Isig (or glucose permeability). Also, the presence of more GLM-acrylate polymer in the polymer blend can enhance the adhesion between the polycarbonate polymeric film layer and the proximal layer in the sensor (eg, the layer comprising glucose oxidase).

Claims (20)

1. A method of increasing the thermal stability of a biocompatible film formed from a reaction mixture, the reaction mixture comprising: diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; a siloxane having amino, hydroxyl or carboxylic acid functional groups at the termini; and a catalyst; The method includes forming the reaction mixture such that the catalyst is present in the reaction mixture in an amount less than 0.2% of the components of the reaction mixture; This increases the thermal energy of the biocompatible membrane compared to a comparable membrane formed from the reaction mixture in which the catalyst is present in the reaction mixture in an amount greater than or equal to 0.2% of the reaction mixture stability. 2. The method of claim 1, wherein the reaction mixture further comprises a polycarbonate diol. 3. The method of claim 1, wherein the catalyst is present in the reaction mixture in an amount less than 0.11% of the reaction mixture. 4. The method of claim 1, wherein: (a) the diisocyanates include hexamethylene diisocyanate and/or methylene diphenyl diisocyanate; (b) the hydrophilic polymer comprising a hydrophilic diol or hydrophilic diamine comprises JEFFAMINE; (c) The siloxanes having amino, hydroxyl or carboxylic acid functional groups at the termini include polydimethylsiloxanes. 5. The method of claim 2, wherein the polycarbonate diol comprises poly(1,6-hexylcarbonate)diol and/or poly(1,6hexyl-1,5amylcarbonate)diol . 6. The biocompatible composition of claim 2, wherein: (a) The diisocyanates include: 17% to 23% by weight of hexamethylene diisocyanate; and 0% to 8.5% by weight of methylene diphenyl diisocyanate; (b) JEFFAMINE includes 28% to 51% by weight of JEFFAMINE 600 and/or JEFFAMINE 900; (c) the polydimethylsiloxane comprises 14% to 48% by weight of Dimethicone-A15; and (d) The polycarbonate diol includes 7.5% to 19% by weight of poly(1,6-hexyl carbonate) diol. 7. The biocompatible composition of claim 6, wherein: (a) The diisocyanates include: about 22% hexamethylene diisocyanate; and about 3.5% methylene diphenyl diisocyanate; (b) JEFFAMINE comprises about 45% JEFFAMINE 600 and/or JEFFAMINE 900; (c) Dimethicone comprising about 22.5% Dimethicone-A15; and (d) The polycarbonate diol includes about 7.5% poly(1,6-hexyl carbonate) diol. 8. The biocompatible composition of claim 1 wherein water is added as a chain extender to the polyurea-polyurethane copolymer reaction mixture. 9. The method of claim 1, wherein the thermal stability of the biocompatible film is determined by observing the biocompatibility for a period of at least 3, 5 or 7 days at a temperature of 60°C The change in molecular weight of the membrane was measured. 10. An amperometric detection analyte sensor comprising: basal layer; a conductive layer disposed on the base layer and including a working electrode; an analyte sensing layer disposed on the conductive layer; and an analyte modulating layer disposed on the analyte sensing layer, wherein the analyte modulating layer: (a) is formed from a reaction mixture comprising: diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; siloxanes having amino, hydroxyl or carboxylic acid functional groups at the termini; and a catalyst, the catalyst being present in the reaction mixture in an amount less than 0.2% of the components of the reaction mixture; and (b) exhibits greater thermal stability than comparable analyte modulating layers formed from the 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. 11. The analyte sensor of claim 10, which is a glucose sensor capable of being implanted in vivo. 12. The analyte sensor of claim 10, further comprising at least one of: a protein layer disposed on the analyte sensing layer; or A cover layer disposed on the analyte sensor device, wherein the cover layer includes apertures positioned on the cover layer to facilitate protection of analytes present in the in vivo environment from contact and diffusion through passing through the analyte modulating layer; and contacting the analyte sensing layer. 13. The analyte sensor of claim 10, wherein the conductive layer comprises a plurality of electrodes including a working electrode, a counter electrode, and a reference electrode. 14. The analyte sensor of claim 13, 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 cells and are positionally distributed on the conductive layer in a repeating pattern of cells. 15. A method of making an analyte sensor for implantation in a mammal, the method comprising the steps of: provide a base layer; forming a conductive layer on the base layer, wherein the conductive layer includes a working electrode; forming an analyte sensing layer on the conductive layer, wherein the analyte sensing layer comprises an oxidoreductase; An analyte modulating layer is formed on the analyte sensing layer, wherein the analyte modulating layer (a) is formed from a reaction mixture comprising: diisocyanate; a hydrophilic polymer comprising a hydrophilic diol or a hydrophilic diamine; siloxanes having amino, hydroxyl or carboxylic acid functional groups at the termini; and a catalyst, the catalyst being present in the reaction mixture in an amount less than 0.2% of the components of the reaction mixture; and (b) exhibits greater thermal stability than comparable analyte modulating layers formed from the 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. 16. The method of claim 15, wherein the reaction mixture further comprises a polycarbonate diol. 17. The method of claim 16, wherein: (a) the diisocyanates include hexamethylene diisocyanate and/or methylene diphenyl diisocyanate; (b) JEFFAMINE comprises about 45% JEFFAMINE 600 and/or JEFFAMINE 900; (c) Dimethicone comprising about 22.5% Dimethicone-A15; and (d) The polycarbonate diol includes about 7.5% poly(1,6-hexyl carbonate) diol. 18. The method of claim 17, wherein: (a) The diisocyanates include: about 22% hexamethylene diisocyanate; and about 3.5% methylene diphenyl diisocyanate; (b) JEFFAMINE comprises about 45% JEFFAMINE 600 and/or JEFFAMINE 900; (c) Dimethicone comprising about 22.5% Dimethicone-A15; and (d) The polycarbonate diol includes about 7.5% poly(1,6-hexyl carbonate) diol. 19. The method of claim 15, wherein water is added as a chain extender to the polyurea-polyurethane copolymer reaction mixture. 20. The method of claim 15, wherein the catalyst is present in the reaction mixture in an amount less than 0.11% of the reaction mixture.

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