US20070227907A1 - Methods and materials for controlling the electrochemistry of analyte sensors - Google Patents

Methods and materials for controlling the electrochemistry of analyte sensors Download PDF

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US20070227907A1
US20070227907A1 US11/397,543 US39754306A US2007227907A1 US 20070227907 A1 US20070227907 A1 US 20070227907A1 US 39754306 A US39754306 A US 39754306A US 2007227907 A1 US2007227907 A1 US 2007227907A1
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Prior art keywords
layer
analyte
electrode
sensor
surface area
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US11/397,543
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Inventor
Rajiv Shah
Udo Hoss
Rebecca K. Gottlieb
Gopikrishnan Soundararajan
James D. Holker
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Medtronic Minimed Inc
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Medtronic Minimed Inc
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Priority to US11/397,543 priority Critical patent/US20070227907A1/en
Priority to US11/492,273 priority patent/US9237865B2/en
Priority to US11/493,054 priority patent/US9492111B2/en
Assigned to MEDTRONIC MINIMED, INC. reassignment MEDTRONIC MINIMED, INC. ASSIGNMENT OF ASSIGNORS INTEREST (SEE DOCUMENT FOR DETAILS). Assignors: HOLKER, JAMES D., SHAH, RAJIV, GOTTLIEB, REBECCA K., SOUNDARARAJAN, GOPIKRISHNAN
Priority to PCT/US2007/008491 priority patent/WO2007114943A2/en
Priority to JP2009504298A priority patent/JP5568298B2/ja
Priority to CA2648151A priority patent/CA2648151C/en
Priority to EP07754927A priority patent/EP2005157A2/en
Publication of US20070227907A1 publication Critical patent/US20070227907A1/en
Priority to US12/644,949 priority patent/US20100096278A1/en
Priority to US12/731,567 priority patent/US9541519B2/en
Abandoned legal-status Critical Current

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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N27/00—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means
    • G01N27/26—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means by investigating electrochemical variables; by using electrolysis or electrophoresis
    • G01N27/28—Electrolytic cell components
    • G01N27/30—Electrodes, e.g. test electrodes; Half-cells
    • G01N27/327—Biochemical electrodes, e.g. electrical or mechanical details for in vitro measurements
    • C—CHEMISTRY; METALLURGY
    • C25—ELECTROLYTIC OR ELECTROPHORETIC PROCESSES; APPARATUS THEREFOR
    • C25D—PROCESSES FOR THE ELECTROLYTIC OR ELECTROPHORETIC PRODUCTION OF COATINGS; ELECTROFORMING; APPARATUS THEREFOR
    • C25D5/00—Electroplating characterised by the process; Pretreatment or after-treatment of workpieces
    • C25D5/10—Electroplating with more than one layer of the same or of different metals
    • C—CHEMISTRY; METALLURGY
    • C25—ELECTROLYTIC OR ELECTROPHORETIC PROCESSES; APPARATUS THEREFOR
    • C25D—PROCESSES FOR THE ELECTROLYTIC OR ELECTROPHORETIC PRODUCTION OF COATINGS; ELECTROFORMING; APPARATUS THEREFOR
    • C25D5/00—Electroplating characterised by the process; Pretreatment or after-treatment of workpieces
    • C25D5/60—Electroplating characterised by the structure or texture of the layers
    • C25D5/605—Surface topography of the layers, e.g. rough, dendritic or nodular layers
    • C—CHEMISTRY; METALLURGY
    • C25—ELECTROLYTIC OR ELECTROPHORETIC PROCESSES; APPARATUS THEREFOR
    • C25D—PROCESSES FOR THE ELECTROLYTIC OR ELECTROPHORETIC PRODUCTION OF COATINGS; ELECTROFORMING; APPARATUS THEREFOR
    • C25D5/00—Electroplating characterised by the process; Pretreatment or after-treatment of workpieces
    • C25D5/60—Electroplating characterised by the structure or texture of the layers
    • C25D5/605—Surface topography of the layers, e.g. rough, dendritic or nodular layers
    • C25D5/611—Smooth layers
    • C—CHEMISTRY; METALLURGY
    • C25—ELECTROLYTIC OR ELECTROPHORETIC PROCESSES; APPARATUS THEREFOR
    • C25D—PROCESSES FOR THE ELECTROLYTIC OR ELECTROPHORETIC PRODUCTION OF COATINGS; ELECTROFORMING; APPARATUS THEREFOR
    • C25D5/00—Electroplating characterised by the process; Pretreatment or after-treatment of workpieces
    • C25D5/60—Electroplating characterised by the structure or texture of the layers
    • C25D5/623—Porosity of the layers
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N27/00—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means
    • G01N27/26—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means by investigating electrochemical variables; by using electrolysis or electrophoresis
    • G01N27/28—Electrolytic cell components
    • G01N27/30—Electrodes, e.g. test electrodes; Half-cells
    • G01N27/307—Disposable laminated or multilayered electrodes
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/145—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue
    • A61B5/1468—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue using chemical or electrochemical methods, e.g. by polarographic means
    • A61B5/1486—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue using chemical or electrochemical methods, e.g. by polarographic means using enzyme electrodes, e.g. with immobilised oxidase
    • C—CHEMISTRY; METALLURGY
    • C25—ELECTROLYTIC OR ELECTROPHORETIC PROCESSES; APPARATUS THEREFOR
    • C25D—PROCESSES FOR THE ELECTROLYTIC OR ELECTROPHORETIC PRODUCTION OF COATINGS; ELECTROFORMING; APPARATUS THEREFOR
    • C25D1/00—Electroforming
    • C25D1/10—Moulds; Masks; Masterforms
    • Y—GENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10—TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10T—TECHNICAL SUBJECTS COVERED BY FORMER US CLASSIFICATION
    • Y10T428/00—Stock material or miscellaneous articles
    • Y10T428/12—All metal or with adjacent metals
    • Y10T428/12389—All metal or with adjacent metals having variation in thickness
    • Y—GENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10—TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10T—TECHNICAL SUBJECTS COVERED BY FORMER US CLASSIFICATION
    • Y10T428/00—Stock material or miscellaneous articles
    • Y10T428/12—All metal or with adjacent metals
    • Y10T428/12479—Porous [e.g., foamed, spongy, cracked, etc.]
    • Y—GENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10—TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10T—TECHNICAL SUBJECTS COVERED BY FORMER US CLASSIFICATION
    • Y10T428/00—Stock material or miscellaneous articles
    • Y10T428/12—All metal or with adjacent metals
    • Y10T428/12493—Composite; i.e., plural, adjacent, spatially distinct metal components [e.g., layers, joint, etc.]
    • Y—GENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10—TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10T—TECHNICAL SUBJECTS COVERED BY FORMER US CLASSIFICATION
    • Y10T428/00—Stock material or miscellaneous articles
    • Y10T428/12—All metal or with adjacent metals
    • Y10T428/12493—Composite; i.e., plural, adjacent, spatially distinct metal components [e.g., layers, joint, etc.]
    • Y10T428/12771—Transition metal-base component
    • Y10T428/12861—Group VIII or IB metal-base component
    • Y10T428/12875—Platinum group metal-base component
    • Y—GENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10—TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10T—TECHNICAL SUBJECTS COVERED BY FORMER US CLASSIFICATION
    • Y10T428/00—Stock material or miscellaneous articles
    • Y10T428/12—All metal or with adjacent metals
    • Y10T428/12493—Composite; i.e., plural, adjacent, spatially distinct metal components [e.g., layers, joint, etc.]
    • Y10T428/12771—Transition metal-base component
    • Y10T428/12861—Group VIII or IB metal-base component
    • Y10T428/12889—Au-base component

Definitions

  • This invention relates to devices and methods for controlling the electrochemistry of analyte sensors such as glucose sensors used in the management of diabetes.
  • Electrochemical measurements are widely used to determine the concentration of specific substances in fluids including biological analytes such as glucose and lactate. Maintaining the appropriate concentrations of glucose in the blood of an individual for example is extremely important for maintaining homeostasis. A concentration of glucose below the normal range, or hypoglycemia, can cause unconsciousness and lowered blood pressure, and may even result in death. A concentration of glucose at levels higher than normal, or hyperglycemia, can result in synthesis of fatty acids and cholesterol, and in diabetics, coma. The measurement of the concentration of glucose in a person's blood, therefore, has become a necessity for diabetics who control the level of blood glucose by insulin therapy.
  • an electrochemical sensor In clinical settings, accurate and relatively fast determinations of glucose and/or lactate levels can be determined from blood samples utilizing electrochemical sensors.
  • the analyte diffuses from the test environment into the sensor housing through a permeable membrane to a working electrode where the analyte chemically reacts.
  • a complementary chemical reaction occurs at a second electrode in the sensor housing known as a counter electrode.
  • the electrochemical sensor produces an analytical signal via the generation of a current arising directly from the oxidation or reduction of the analyte at the working and counter electrodes.
  • an electrolytic electrochemical sensor often includes a third electrode, commonly referred to as a reference electrode.
  • a reference electrode is used to maintain the working electrode at a known voltage or potential.
  • the electrodes of an electrochemical sensor provide a surface at which an oxidation or a reduction reaction occurs (that is, an electrochemically active surface) to provide a mechanism whereby the ionic conduction of an electrolyte solution in contact with the electrodes is coupled with the electron conduction of each electrode to provide a complete circuit for a current.
  • the electrode at which an oxidation occurs is the anode
  • the electrode at which the “complementary” reduction occurs is the cathode.
  • the working and counter electrode in combination produce an electrical signal that is both related to the concentration of the analyte and is sufficiently strong to provide a signal-to-noise ratio suitable to distinguish between concentration levels of the analyte over the entire range of interest.
  • a common type of glucose or lactate electrode sensor comprises an enzyme electrode which utilizes an enzyme to convert glucose or lactate to an electroactive product which is then analyzed electrochemically.
  • a chemical reaction at the electrode converts glucose in the presence of enzymes, such as glucose oxidase, and results in the formation of reaction products including hydrogen peroxide.
  • glucose reacts with oxygen to form gluconolactone and hydrogen peroxide.
  • a suitable electrode can then measure the formation of hydrogen peroxide as an electrical signal. The signal is produced following the transfer of electrons from the peroxide to the electrode, and under suitable conditions, the enzyme catalyzed flow of current is proportional to the glucose concentration.
  • Lactate electrode sensors including an enzyme electrode similarly convert lactate in the presence of enzymes, such as lactate oxidase.
  • glucose and oxygen from blood diffuse through a primary membrane of the sensor.
  • an enzyme such as glucose oxidase, catalyzes the conversion of glucose to hydrogen peroxide and gluconolactone.
  • the hydrogen peroxide may diffuse back through the primary membrane, or it may further diffuse through the secondary membrane to an electrode where it can be reacted to form oxygen and a proton to produce a current proportional to the glucose concentration.
  • Embodiments of the invention disclosed herein include electrochemical analyte sensors comprising elements such as electrodes and/or electrode combinations (e.g. working and counter electrode combinations) designed to optimize factors including the reactivity, sensitivity, functioning and lifespan of the analyte sensors.
  • An illustrative embodiment of the invention is a method of performing an electrochemical reaction within an analyte sensor comprising using an analyte sensor constructed to include an electrode layer configuration that is designed to optimize the electrochemical reaction at the electrode when the electrode is exposed to an analyte.
  • the analyte sensor typically includes at least one electrode disposed upon a base substrate where this base substrate comprises a geometric feature selected to increase the surface area of an electrochemically reactive surface of the electrode disposed thereon such that surface area-to-volume ratio of the electrochemically-reactive surface area of the electrode disposed on the geometric feature is greater than surface-area-to-volume ratio of the reactive surface of the electrode when disposed on a flat surface.
  • the surface area to volume ratio of the electrochemically reactive surface area of the electrode disposed on the geometric feature is at least 10%, 25%, 50%, 75% or 100% greater than surface area-to-volume ratio of the reactive surface of the electrode when disposed on a flat surface.
  • an electrode in the analyte sensor comprises a porous matrix.
  • the porous matrix can have a surface area that is at least 2, 4, 6, 8, 10, 12, 14, 16 or 18 times the surface area of an essentially non-porous matrix of same dimensions.
  • Analyte sensors having electrodes constructed to have this type of configuration can be constructed by a variety of methods known in the art; for example, by disposing the electrode material (e.g. a metal such as platinum) on a base substrate adjoining layer that includes a geometric feature comprising a lip, a shoulder, a ridge, a notch, a depression, a channel or the like.
  • the geometric feature of the base substrate causes the electrochemically reactive surface area of the electrode to form a nodular geometry or the like.
  • the analyte sensor comprises a plurality of discrete geometric features having a plurality of electrochemically reactive electrode surfaces.
  • Such pluralities of features can include patterns such as rows of depressions and/or ridges or the like; for example, a row of ridges resembling zebra stripes.
  • the analyte sensor comprises at least 2, 3, 4, 5, 6, 7, 8, 9 or 10 discrete geometric features having a plurality of electrochemically reactive electrode surfaces.
  • analyte sensor further comprises an analyte sensing layer disposed on the electrode having a relatively high surface area-to-volume ratio of the reactive surface, wherein the analyte sensing layer detectably alters the electrical current at the electrode in the presence of an analyte; an optional protein layer disposed on the analyte sensing layer; an adhesion promoting layer disposed on the analyte sensing layer or the optional protein layer, wherein the adhesion promoting layer promotes the adhesion between the analyte sensing layer and an analyte modulating layer disposed on the analyte sensing layer; an analyte modulating layer disposed on the analyte sensing layer, wherein the analyte modulating layer modulates the diffusion of the analyte therethrough; and an optional cover layer disposed on at least a portion of the analyte modulating layer, wherein the cover layer further includes an aperture over at least a
  • the analyte sensor is designed to be implantable within a mammal.
  • the electrochemical reaction at the electrode involves a protein reactive with an analyte present in mammalian blood such as glucose oxidase, glucose dehydrogenase, lactate oxidase, hexokinase or lactate dehydrogenase.
  • a related embodiment of the invention is an analyte sensor for detecting an analyte in a fluid, the apparatus comprising at least one electrode disposed upon a base substrate, wherein the base substrate includes a geometric feature selected to increase the surface area of an electrochemically reactive surface on the electrode deposited thereon (e.g. a lip, a shoulder, a ridge, a notch, a depression, a channel or the like) such that surface-area-to-volume ratio of the electrochemically reactive surface area of the electrode disposed on the geometric feature is greater than surface area-to-volume ratio of the reactive surface of the electrode when disposed on a flat surface.
  • a geometric feature selected to increase the surface area of an electrochemically reactive surface on the electrode deposited thereon (e.g. a lip, a shoulder, a ridge, a notch, a depression, a channel or the like) such that surface-area-to-volume ratio of the electrochemically reactive surface area of the electrode disposed on the geometric feature is greater than surface area-
  • the analyte sensor comprises a plurality of discrete geometric features having a plurality of electrochemically reactive electrode surfaces.
  • the analyte sensor is implantable and comprises an analyte sensing layer disposed on the electrode, wherein the analyte sensing layer detectably alters the electrical current at the electrode in the presence of an analyte; an optional protein layer disposed on the analyte sensing layer; an adhesion promoting layer disposed on the analyte sensing layer or the optional protein layer, wherein the adhesion promoting layer promotes the adhesion between the analyte sensing layer and an analyte modulating layer disposed on the analyte sensing layer; and an analyte modulating layer disposed on the analyte sensing layer, wherein the analyte modulating layer modulates the diffusion of the analyte therethrough; and an optional cover layer disposed on at least a portion of the analyte modulating
  • Yet another embodiment of the invention is a method of modulating electrochemical reactions within an implantable analyte sensor, the method comprising performing electrochemical reactions within an implantable analyte sensor comprising a working electrode having a reactive surface area, wherein during analyte sensing, the working electrode generates electrons that reduce a plurality of composition species in the electrochemical reaction including oxygen (O 2 ); and a counter electrode having a reactive surface area, wherein the size of the reactive surface area of the counter electrode is selected so as to control the reduction of the plurality of composition species in the electrochemical reaction so that oxygen (O 2 ) is the predominant composition species reduced by the electrons generated at the working electrode, wherein the compound substrates having a first affinity for the electrons exhibit an affinity that is higher than the affinity of the compound substrates having a second affinity for the electrons; and a counter electrode of a second surface area, the second surface area is selected to be a size that reduces the interaction between electrons generated at the working electrode with compound substrates having the second affinity for the electrons generated
  • a related embodiment of the invention is an implantable electrochemical analyte sensor designed to include this electrode architecture.
  • the working electrode and the counter electrode in the analyte sensor comprise a porous matrix.
  • the working electrode can comprise a relatively nonporous matrix while the counter electrode can comprise a porous matrix or vice versa.
  • the porous matrix has a surface area that is at least 1.5, 2, 2.5, 3, 3.5, 4, 4.5, 5, 5.5, 6, 6.5, 7, 7.5, 8, 10, 12, 14, 16 or 18 times the surface area of an essentially non-porous matrix of same dimensions.
  • the implantable analyte sensor further comprises an analyte sensing layer disposed on the working electrode, wherein the analyte sensing layer detectably alters the electrical current at the working electrode in the presence of an analyte; an optional protein layer disposed on the analyte sensing layer; an adhesion promoting layer disposed on the analyte sensing layer or the optional protein layer, wherein the adhesion promoting layer promotes the adhesion between the analyte sensing layer and an analyte modulating layer disposed on the analyte sensing layer; and an analyte modulating layer disposed on the analyte sensing layer, wherein the analyte modulating layer modulates the diffusion of the analyte therethrough; and an optional cover layer disposed on at least a portion of the analyte modulating layer, wherein the cover layer further includes an aperture over at least a portion of the analyte modulating layer.
  • Yet another illustrative embodiment of the invention is a method of making a metallic electrode by electrodepositing a plurality of metal layers that comprise the electrode using cycles of differing electroplating conditions.
  • the method comprises a first cycle of electroplating where a metal is electrodeposited onto a substrate under a first set of conditions selected to produce a first metal layer having a first surface area and a first adhesion strength between the substrate and the first metal layer.
  • the method then involves a second cycle of electroplating where a metal composition is then electrodeposited onto the first metal layer under a second set of conditions selected to produce a second metal layer having a second surface area and a second adhesion strength between the first metal layer and the second metal layer.
  • the first and second set of conditions are selected to produce a second metal layer having a second surface area that is greater than the first surface area of the first metal layer produced by the first set of conditions and a second metal layer having an adhesion with the first metal layer that is greater than the adhesion between the first metal layer and the substrate produced by the first set of conditions.
  • the method further comprises additional cycles of electroplating.
  • the method comprises a third cycle of electroplating where a metal composition is electrodeposited onto the second layer under a third set of conditions selected to produce a third metal layer having a third surface area.
  • the second and third set of conditions are selected to produce a third metal layer having a greater density than the density of the second metal layer.
  • the third set of conditions produces a third metal layer having a third surface area that is less than the second surface area of the second metal layer produced by the second set of conditions.
  • the invention also provides additional articles of manufacture including sensor elements, sensor sets and kits.
  • a kit and/or sensor element or set useful for the sensing an analyte as is described above, is provided.
  • the kit and/or sensor set typically comprises a container, a label and a sensor as described above.
  • the typical embodiment is a kit comprising a container and, within the container, an analyte sensor apparatus having a design as disclosed herein and instructions for using the analyte sensor apparatus.
  • FIG. 1 provides a schematic of the well known reaction between glucose and glucose oxidase.
  • this reaction involves glucose oxidase (GOx), glucose and oxygen in water.
  • glucose oxidase GOx
  • glucose and oxygen in water.
  • the enzyme In the reductive half of the reaction, two protons and electrons are transferred from ⁇ -D-glucose to the enzyme yielding d-gluconolactone.
  • the enzyme is oxidized by molecular oxygen yielding hydrogen peroxide.
  • the d-gluconolactone then reacts with water to hydrolyze the lactone ring and produce gluconic acid.
  • the hydrogen peroxide produced by this reaction is oxidized at the working electrode (H 2 O 2 ⁇ 2H+ + O 2 + 2e ⁇ ).
  • FIG. 2 provides a diagrammatic view of a typical analyte sensor configuration of the current invention.
  • analyte as used herein is a broad term and is used in its ordinary sense, including, without limitation, to refer to a substance or chemical constituent in a fluid such as a biological fluid (for example, blood, interstitial fluid, cerebral spinal fluid, lymph fluid or urine) that can be analyzed.
  • a biological fluid for example, blood, interstitial fluid, cerebral spinal fluid, lymph fluid or urine
  • Analytes can include naturally occurring substances, artificial substances, metabolites, and/or reaction products.
  • the analyte for measurement by the sensing regions, devices, and methods is glucose.
  • other analytes are contemplated as well, including but not limited to, lactate. Salts, sugars, proteins fats, vitamins and hormones naturally occurring in blood or interstitial fluids can constitute analytes in certain embodiments.
  • the analyte can be naturally present in the biological fluid or endogenous; for example, a metabolic product, a hormone, an antigen, an antibody, and the like.
  • the analyte can be introduced into the body or exogenous, for example, a contrast agent for imaging, a radioisotope, a chemical agent, a fluorocarbon-based synthetic blood, or a drug or pharmaceutical composition, including but not limited to insulin.
  • a contrast agent for imaging for example, a radioisotope, a chemical agent, a fluorocarbon-based synthetic blood, or a drug or pharmaceutical composition, including but not limited to insulin.
  • the metabolic products of drugs and pharmaceutical compositions are also contemplated analytes.
  • the term “sensor,” as used herein, is a broad term and is used in its ordinary sense, including, without limitation, the portion or portions of an analyte-monitoring device that detects an analyte.
  • the sensor includes an electrochemical cell that has a working electrode, a reference electrode, and optionally a counter electrode passing through and secured within the sensor body forming an electrochemically reactive surface at one location on the body, an electronic connection at another location on the body, and a membrane system affixed to the body and covering the electrochemically reactive surface.
  • a biological sample for example, blood or interstitial fluid
  • an enzyme for example, glucose oxidase
  • electrochemical cell is a broad term and is used in its ordinary sense, including, without limitation, a device in which chemical energy is converted to electrical energy.
  • a cell typically consists of two or more electrodes held apart from each other and in contact with an electrolyte solution. Connection of the electrodes to a source of direct electric current renders one of them negatively charged and the other positively charged. Positive ions in the electrolyte migrate to the negative electrode (cathode) and there combine with one or more electrons, losing part or all of their charge and becoming new ions having lower charge or neutral atoms or molecules; at the same time, negative ions migrate to the positive electrode (anode) and transfer one or more electrons to it, also becoming new ions or neutral particles. The overall effect of the two processes is the transfer of electrons from the negative ions to the positive ions, a chemical reaction.
  • electrochemically reactive surface and “electroactive surface” as used herein are broad terms and are used in their ordinary sense, including, without limitation, the surface of an electrode where an electrochemical reaction takes place.
  • a working electrode measures hydrogen peroxide produced by the enzyme catalyzed reaction of the analyte being detected reacts creating an electric current (for example, detection of glucose analyte utilizing glucose oxidase produces H 2 O 2 as a by product, H 2 O 2 reacts with the surface of the working electrode producing two protons (2H + ), two electrons (2e ⁇ ) and one molecule of oxygen (O 2 ) which produces the electronic current being detected).
  • a reducible species for example, O 2 is reduced at the electrode surface in order to balance the current being generated by the working electrode.
  • sensing region is a broad term and is used in its ordinary sense, including, without limitation, the region of a monitoring device responsible for the detection of a particular analyte.
  • the sensing region can comprise a non-conductive body, a working electrode, a reference electrode, and a counter electrode passing through and secured within the body forming electrochemically reactive surfaces on the body and an electronic connective means at another location on the body, and a one or more layers covering the electrochemically reactive surface.
  • electrical potential and “potential” as used herein, are broad terms and are used in their ordinary sense, including, without limitation, the electrical potential difference between two points in a circuit which is the cause of the flow of a current.
  • system noise is a broad term and is used in its ordinary sense, including, without limitation, unwanted electronic or diffusion-related noise which can include Gaussian, motion-related, flicker, kinetic, or other white noise, for example.
  • interferants and “interfering species,” as used herein, are broad terms and are used in their ordinary sense, including, but not limited to, effects and/or species that interfere with the measurement of an analyte of interest in a sensor to produce a signal that does not accurately represent the analyte measurement.
  • interfering species are compounds with an oxidation potential that overlaps with the analyte to be measured.
  • embodiments of the invention relate to the use of an electrochemical sensor that measures a concentration of an analyte of interest or a substance indicative of the concentration or presence of the analyte in fluid.
  • the sensor is a continuous device, for example a subcutaneous, transdermal, or intravascular device.
  • the device can analyze a plurality of 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.
  • the senor is of the type that senses a product or reactant 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 typically comprise a membrane surrounding the enzyme through which a bodily fluid passes and in which an analyte within the bodily fluid reacts with an enzyme in the presence of oxygen to generate a product.
  • the product is then measured using electrochemical methods and thus the output of an electrode system functions as a measure of the analyte.
  • the sensor can use an amperometric, coulometric, conductimetric, and/or potentiometric technique for measuring the analyte.
  • Embodiments of the invention described herein can be adapted and implemented with a wide variety of known electrochemical sensors, including for example, U.S. Patent Application No. 20050115832, U.S. Pat. Nos. 6,001,067, 6,702,857, 6,212,416, 6,119,028, 6,400,974, 6,595,919, 6,141,573, 6,122,536, 6,512,939 5,605,152, 4,431,004, 4,703,756, 6,514,718, 5,985,129, 5,390,691, 5,391, 250, 5,482,473, 5,299,571, 5,568,806, 5,494,562, 6,120,676, 6,542,765 as well as PCT International Publication Numbers 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
  • Embodiments of the invention disclosed herein provide sensors of the type used, for example, in subcutaneous or transcutaneous monitoring of blood glucose levels in a diabetic patient.
  • a variety of implantable, electrochemical biosensors have been developed for the treatment of diabetes and other life-threatening diseases.
  • Many existing sensor designs use some form of immobilized enzyme to achieve their bio-specificity.
  • a first class of glucose sensor designs use a very thin ( ⁇ 1 micron) layer of glucose oxidase (GOx) and bovine serum albumin that is either spray or spin coated onto the working electrode and cross-linked with glutaraldehyde.
  • a second class of glucose sensor design employs a thick ( ⁇ 1 mm) hydrogel known as the Sensor Matrix Protein (SMP), which typically consists of an enzyme such as GOx and human serum albumin cross-linked together with a cross-linking agent such as glutaraldehyde.
  • SMP Sensor Matrix Protein
  • the immobilized enzyme configurations of the two above-noted classes of sensor designs possess different advantages that serve to increase operational sensor life. Due to the close proximity of the immobilized GOx to the peroxide-consuming electrode, the first class of sensor designs are believed to possess significantly decreased enzyme deactivation rate constants.
  • the thick SMPs utilized in the second class of sensor designs can incorporate orders of magnitude more enzyme than the first class.
  • a matrix such as an enzymatic hydrogel matrix to function.
  • matrix is used herein according to its art-accepted meaning of something within or from which something else originates, develops, takes form and/or is found.
  • An exemplary enzymatic hydrogel matrix for example typically comprises a bio-sensing enzyme (e.g. glucose oxidase or lactate oxidase) and human serum albumin proteins that have been cross-linked together with a crosslinking agent such as glutaraldehyde to form a polymer network. This network is then swollen with an aqueous solution to form an enzymatic hydrogel matrix.
  • bio-sensing enzyme e.g. glucose oxidase or lactate oxidase
  • human serum albumin proteins that have been cross-linked together with a crosslinking agent such as glutaraldehyde to form a polymer network. This network is then swollen with an aqueous solution to form an enzymatic hydrogel matrix.
  • Embodiments of the invention disclosed herein provide sensor elements having enhanced material properties and sensors constructed from such elements.
  • the disclosure further provides methods for making and using such sensors.
  • a variety of the elements disclosed herein e.g. electrodes and electrode designs
  • the analyte sensor elements, architectures and methods for making and using these elements that are disclosed herein can be used to establish a variety of layered sensor structures.
  • Such sensors of the invention exhibit a surprising degree of flexibility and versatility, characteristic which allow a wide variety of sensor configurations to be designed to examine a wide variety of analyte species.
  • the transduction of the analyte concentration into a processable signal is by electrochemical means.
  • These transducers may include any of a wide variety of amperometric, potentiometric, or conductimetric base sensors known in the art.
  • the microfabrication sensor techniques and materials of the instant invention may be applied to other types of transducers (e.g., acoustic wave sensing devices, thermistors, gas-sensing electrodes, field-effect transistors, optical and evanescent field wave guides, and the like) fabricated in a substantially nonplanar, or alternatively, a substantially planar manner.
  • Embodiments of the sensors disclosed herein incorporate one or more sensor elements having enhanced material properties.
  • Embodiments of the invention include sensor including these elements and well methods for making and using them.
  • Embodiments of the invention disclosed herein include electrochemical analyte sensors comprising elements such as electrodes and/or electrode combinations (e.g. working an counter electrode combinations) designed to optimize factors including the reactivity, sensitivity, functioning and lifespan of the analyte sensors.
  • Certain specific embodiments of the invention are designed to optimize the electrochemical reactions that function in the sensing of an analyte of interest.
  • the optimized embodiments of the invention disclosed herein can be utilized and/or applied to a wide variety of sensor methods and designs. The following sections describe illustrative sensor elements, sensor configurations and methodological embodiments of the invention.
  • embodiments of the invention are designed to enhance sensor stability.
  • embodiments of the invention include analyte sensors having a plurality of adjoining layers consisting of different functional constituents.
  • one or more of the layers can become unstable and separate in part or in whole from an adjoining layer.
  • Such delamination events can compromise the function of one or more of the different constituents and consequently have a negative impact on a sensor's function, for example its long term stability.
  • a series of flat, uniform layers typically have better adherence between layers and are therefore more stable than sensors having disparate layers of rough and/or polymorphic constituents.
  • layer that can exhibit problems with delamination is an electrode layer that is deposited upon a base layer in a manner that creates an irregular surface, for example when a platinum electrode layer is deposited on a base substrate that has an edge or shoulder that causes a nodule of a electrode material (e.g. platinum black) to form over this feature on the base substrate.
  • a electrode material e.g. platinum black
  • This can happen, for example, when platinum black is disposed in the sensor in a manner that allows it to grow up the side of a material already in the sensor; for example, an edge of an insulating composition.
  • one embodiment of the invention is a method of contributing to the uniformity of a layer in an analyte sensing device comprising a plurality of layers, the method comprising disposing an electrode used in the device on a base substrate, wherein the base substrate is selected to exhibit an essentially or predominantly flat geometry so that the electrode disposed on the base substrate also exhibits a relatively flat geometry, thereby contributing to the uniformity of a layer in a analyte sensing device.
  • the sensor can be manufactured according to a process that avoids depositing the electrode materials on geometric features already present in the sensor.
  • the base substrate is selected to exhibit a flat geometry by avoiding disposing the electrode on a ridge, lip, shoulder or edge generated by an insulating composition that is disposed in an area where the electrode is disposed.
  • a related embodiment of the invention is a method of contributing to the uniformity of a layer in an analyte sensing device comprising a plurality of layers, the method including disposing an electrode used in the device on a base substrate, and then removing a portion of the electrode so disposed on the base substrate, wherein the portion of the electrode disposed on the base substrate that is removed is an irregular feature such as a bulge or the like that results from the electrode being disposed on a lip, shoulder, edge or the like etc. that is present in another of the plurality of layers that is disposed in an area where the electrode is disposed.
  • the resulting layer having the electrode where the bulge or nodule is removed exhibits a more uniform surface than does the layer where the bulge is not removed.
  • Such bulges can be removed by one of a variety of methods known in the art, for example, via mechanical or chemical processes.
  • these methods of contributing to the uniformity of one or more layers in an analyte sensor have the result of inhibiting delamination of a layer within the plurality of layers.
  • These methods of contributing to the uniformity of one or more layers in an analyte sensor can also increase sensor stability; for example, the in vivo stability of a sensor.
  • sensors having electrode surfaces with unusual and/or non-uniform geometries have unexpected and desirable properties.
  • the electrode material e.g. platinum
  • the beneficial properties of these shaped electrodes outweigh the detrimental effects that result from a lack of uniformity in the electrode layer. This unexpected observation forms the basis of certain embodiments of the invention discussed below.
  • One illustrative embodiment of the invention is a method of performing an electrochemical reaction within an analyte sensor comprising using an analyte sensor constructed to include an electrode layer configuration that is designed to optimize the electrochemical reaction at the electrode when the electrode is exposed to an analyte.
  • the analyte sensor typically includes at least one electrode disposed upon a base substrate where this base substrate comprises a geometric feature selected to increase the surface area of an electrochemically reactive surface of the electrode disposed thereon such that the surface area-to-volume ratio of the electrochemically reactive surface area of the electrode disposed on the geometric feature is greater than surface area-to-volume ratio of the reactive surface of the electrode when disposed on a flat surface.
  • the electrode further comprises a porous matrix.
  • the porous matrix has a surface area that is at least 2, 4, 6, 8, 10, 12, 14, 16 or 18 times the surface area of an essentially non-porous matrix of same dimensions.
  • the electronic signal in response to exposure to the analyte that is generated at the electrochemically reactive surface area of the electrode disposed on the geometric feature is greater than the electronic signal generated an electrochemically reactive surface area of the electrode when the electrode is disposed on a flat surface.
  • the in vivo lifetime of the analyte sensor having the electrochemically reactive surface area of the electrode disposed on the geometric feature is greater than the in vivo lifetime of an analyte sensor having an electronic signal generated on an electrochemically reactive surface area of an electrode when the electrode is disposed on a flat surface.
  • the cerebral cortex of the human brain is highly convoluted, meaning it has many folds and creases, convolutions that allow a large surface area of brain (and an associated larger number of neurons) to fit inside our relatively small skulls.
  • the surface area to volume ratio of the electrochemically reactive surface area of the electrode disposed on the geometric feature is at least 10%, 25%, 50%, 75% or 100% greater than surface area-to-volume ratio of the reactive surface of the electrode when disposed on a flat surface.
  • Analyte sensor having electrodes constructed to have this type of configuration can be constructed by a variety of methods known in the art, for example by disposing the electrode material (e.g. a metal such as platinum) on a base substrate adjoining layer that includes a geometric feature comprising a lip, a shoulder, a ridge, a notch, a depression, a channel or the like.
  • the geometric feature of the base substrate causes the electrochemically reactive surface area of the electrode to form a nodule or the like.
  • electrodes having these properties can be generated by other processes known in the art.
  • electrodes having a high surface area to volume ratio can be premade and then subsequently disposed within the analyte sensor.
  • the analyte sensor comprises a plurality of discrete geometric features having a plurality of electrochemically reactive electrode surfaces.
  • Such pluralities of features can include patterns such as rows of depressions and/or ridges or the like, for example a row of ridges resembling zebra stripes.
  • the sensor is manufactured to include these pluralities of features on which the electrode material can be deposited.
  • the analyte sensor comprises at least 2, 3, 4, 5, 6, 7, 8, 9 or 10 discrete geometric features having a plurality of electrochemically reactive electrode surfaces.
  • analyte sensor further comprises an analyte sensing layer disposed on the electrode having a relatively high surface area-to-volume ratio of the reactive surface, wherein the analyte sensing layer detectably alters the electrical current at the electrode in the presence of an analyte; an optional protein layer disposed on the analyte sensing layer; an adhesion promoting layer disposed on the analyte sensing layer or the optional protein layer, wherein the adhesion promoting layer promotes the adhesion between the analyte sensing layer and an analyte modulating layer disposed on the analyte sensing layer; and an analyte modulating layer disposed on the analyte sensing layer, wherein the analyte modulating layer modulates the diffusion of the analyte therethrough; and an optional cover layer disposed on at least a portion of the analyte modulating layer, wherein the cover layer further includes an aperture over at least
  • the analyte sensor is designed to be implantable within a mammal.
  • the electrochemical reaction at the electrode involves a protein reactive with an analyte present in mammalian blood such as glucose oxidase, glucose dehydrogenase, lactate oxidase, hexokinase or lactate dehydrogenase.
  • an analyte present in mammalian blood such as glucose oxidase, glucose dehydrogenase, lactate oxidase, hexokinase or lactate dehydrogenase.
  • hydrogen peroxide is oxidized at the electrochemically reactive surface area of the electrode disposed on the geometric feature of the base substrate.
  • a related embodiment of the invention is an analyte sensor for detecting an analyte in a fluid, the apparatus comprising at least one electrode disposed upon a base substrate, wherein the base substrate includes a geometric feature selected to increase the surface area of an electrochemically reactive surface on the electrode deposited thereon (e.g. a lip, a shoulder, a ridge, a notch, a depression, a channel or the like) such that surface area-to-volume ratio of the electrochemically reactive surface area of the electrode disposed on the geometric feature is greater than surface area-to-volume ratio of the reactive surface of the electrode when disposed on a flat surface.
  • a geometric feature selected to increase the surface area of an electrochemically reactive surface on the electrode deposited thereon (e.g. a lip, a shoulder, a ridge, a notch, a depression, a channel or the like) such that surface area-to-volume ratio of the electrochemically reactive surface area of the electrode disposed on the geometric feature is greater than surface area-to-
  • the analyte sensor comprises a plurality of discrete geometric features having a plurality of electrochemically reactive electrode surfaces.
  • the analyte sensor is implantable and comprises an analyte sensing layer disposed on the electrode, wherein the analyte sensing layer detectably alters the electrical current at the electrode in the presence of an analyte; an optional protein layer disposed on the analyte sensing layer; an adhesion promoting layer disposed on the analyte sensing layer or the optional protein layer, wherein the adhesion promoting layer promotes the adhesion between the analyte sensing layer and an analyte modulating layer disposed on the analyte sensing layer; and an analyte modulating layer disposed on the analyte sensing layer, wherein the analyte modulating layer modulates the diffusion of the analyte therethrough; and an optional cover layer disposed on at least a portion of the analyte modulating
  • Yet another embodiment of the invention is a method of modulating electrochemical reactions within an implantable analyte sensor, the method comprising performing electrochemical reactions within an implantable analyte sensor comprising: a working electrode having a first surface area, wherein electrochemical reactions at the working electrode generate electrons that interact with compound substrates having a first affinity for the electrons that is greater than or equal to that of oxygen (O 2 ), and compound substrates having a second affinity for the electrons that is less than or equal to that of oxygen (O 2 ), wherein the second surface area is selected to be a size that reduces the interaction between electrons generated at the working electrode with compound substrates having the second affinity for the electrons generated at the working electrode; so that electrochemical reactions within the implantable analyte sensor are modulated.
  • a closely related embodiment of the invention is a method of modulating electrochemical reactions within an implantable analyte sensor, the method comprising performing electrochemical reactions within an implantable analyte sensor comprising: a working electrode having a reactive surface area, wherein during analyte sensing, the working electrode generates electrons that reduce a plurality of composition species in the electrochemical reaction including oxygen (O 2 ); and a counter electrode having a reactive surface area, wherein the size of the reactive surface area of the counter electrode (e.g.
  • the term predominant is used herein according to its art accepted meaning of being the most frequent or common. In one illustrative example of such an embodiment, 51% of the electrons generated at the working electrode reduce oxygen (O 2 ) species over the lifetime of the sensor.
  • the methods (and associated sensor architectures) designed to selectively reduce oxygen (O 2 ) species in the electrochemical reactions have a number of surprising advantages over existing embodiments in the art.
  • the sensing of the analyte is therefore occurs in a controlled environment.
  • this sensor structure maintains a greater consistency in the electrochemical reaction environment (leading to a greater consistency in signals generated by the electrochemical reactions) during the period that the analyte sensor is in use.
  • analyte sensors employ a oxidase in analyte sensing such as the protein glucose oxidase
  • using this sensor structure to control the electrochemical reaction so that oxygen (O 2 ) is the predominant composition species reduced by the electrons generated at the working electrode over the functional lifetime of the sensor provides a benefit by providing a stabilized reactive environment for the electrochemical reactions that function to provide the sensor signal.
  • analyte sensors having working and counter electrodes of different sizes may be described in the art, the art fails to teach or suggest the invention described herein, e.g. sensors constructed to include elements specifically designed to control the electrochemical reactions in the described manner (i.e. so that oxygen (O 2 ) is the predominant composition species reduced by electrons generated at the working electrode).
  • embodiments of the sensors disclosed herein are designed to vary the size of the reactive surface area of the counter electrode relative to the size of the reactive surface area of the working electrode, and not just the size of the electrodes.
  • the counter and working electrodes can be of approximately the same size.
  • the counter electrode can be constructed to exhibit an architecture that provides a greater surface area/size ratio (e.g.
  • the greater size of the reactive surface area of the counter electrode e.g. relative to the size of the reactive surface area of the working electrode
  • oxygen (O 2 ) is the predominant composition species reduced by the electrons generated at the working electrode
  • the working electrode and the counter electrode in analyte sensors having the above-noted architectures comprise a porous matrix.
  • the working electrode comprises a relatively nonporous matrix while the counter electrode comprises a porous matrix or vice versa.
  • the porous matrix has a surface area that is at least 2, 4, 6, 8, 10, 12, 14, 16 or 18 times the surface area of an essentially non-porous matrix of same dimensions. In such methods the surface area of the counter electrode is typically about 1.5, 2, 2.5 or 3 times the size of the working electrode.
  • the implantable analyte sensor comprises an analyte sensing layer disposed on the working electrode, wherein the analyte sensing layer detectably alters the electrical current at the working electrode in the presence of an analyte; an optional protein layer disposed on the analyte sensing layer; an adhesion promoting layer disposed on the analyte sensing layer or the optional protein layer, wherein the adhesion promoting layer promotes the adhesion between the analyte sensing layer and an analyte modulating layer disposed on the analyte sensing layer; and an analyte modulating layer disposed on the analyte sensing layer, wherein the analyte modulating layer modulates the diffusion of the analyte therethrough; and an optional cover layer disposed on at least a portion of the analyte modulating layer, wherein the cover layer further includes an aperture over at least a portion of the analyte modulating layer.
  • a related embodiment of the invention is an implantable electrochemical analyte sensor comprising a working electrode having a first surface area, wherein electrochemical reactions at the working electrode generate electrons that interact with compound substrates having a first affinity for the electrons and compound substrates having a second affinity for the electrons, wherein the compound substrates having a first affinity for the electrons exhibit an affinity that is higher than the affinity of the compound substrates having a second affinity for the electrons; and a counter electrode of a second surface area, the second surface area is selected to be a size that reduces the interaction between electrons generated at the working electrode with compound substrates having the second affinity for the electrons generated at the working electrode; an analyte sensing layer disposed on the working electrode, wherein the analyte sensing layer detectably alters the electrical current at the working electrode in the presence of an analyte; an adhesion promoting layer disposed on the analyte sensing layer or the protein layer, wherein the adhesion promoting layer promotes the adhesion between
  • FIG. 2 illustrates a cross-section of a typical sensor structure 100 of the present invention.
  • the sensor is formed from a plurality of components that are typically in the form of layers of various conductive and non-conductive constituents disposed on each other according to a method of the invention to produce a sensor structure.
  • the components of the sensor are typically characterized herein as layers because, for example, it allows for a facile characterization of the sensor structure shown in FIG. 2 .
  • the sensor constituents are combined such that multiple constituents form one or more heterogeneous layers.
  • those of skill in the art understand that the ordering of the layered constituents can be altered in various embodiments of the invention.
  • the embodiment shown in FIG. 2 includes a base layer 102 to support the sensor 100 .
  • the base layer 102 can be made of a material such as a metal and/or a ceramic and/or a polymeric substrate, which may be self-supporting or further supported by another material as is known in the art.
  • Embodiments of the invention include a conductive layer 104 which is disposed on and/or combined with the base layer 102 .
  • the base layer 102 and/or the conductive layer 104 can be constructed to produce electrodes having a configuration where the electrochemically reactive surface area of an electrode is disposed on the geometric feature so that the electrochemically reactive surface area is greater than if it was disposed on a flat surface.
  • the conductive layer 104 comprises one or more electrodes.
  • An operating 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 an electrode that performs multiple functions, for example one that functions as both as a reference and a counter electrode. Still other embodiments may utilize a separate reference element not formed on the sensor. Typically these electrodes are electrically isolated from each other, while situated in close proximity to one another.
  • the base layer 102 and/or conductive layer 104 can be generated using many known techniques and materials.
  • the electrical circuit of the sensor is defined by etching the disposed conductive layer 104 into a desired pattern of conductive paths.
  • a typical electrical circuit for the sensor 100 comprises two or more adjacent conductive paths with regions at a proximal end to form contact pads and regions at a distal end to form sensor electrodes.
  • An electrically insulating cover layer 106 such as a polymer coating is optionally disposed on portions of the sensor 100 .
  • Acceptable polymer coatings for use as the insulating protective cover layer 106 can include, but are not limited to, non-toxic biocompatible polymers such as silicone compounds, polyimides, biocompatible solder masks, epoxy acrylate copolymers, or the like.
  • one or more exposed regions or apertures 108 can be made through the cover layer 106 to open the conductive layer 104 to the external environment and to, for example, allow an analyte such as glucose to permeate the layers of the sensor and be sensed by the sensing elements.
  • Apertures 108 can be formed by a number of techniques, including laser ablation, tape masking, chemical milling or etching or photolithographic development or the like.
  • a secondary photoresist can also be applied to the protective layer 106 to define the regions of the protective layer to be removed to form the aperture(s) 108 .
  • the exposed electrodes and/or contact pads can also undergo secondary processing (e.g. through the apertures 108 ), such as additional plating processing, to prepare the surfaces and/or strengthen the conductive regions.
  • an analyte sensing layer 110 (which is typically a sensor chemistry layer, meaning that materials in this layer undergo a chemical reaction to produce a signal that can be sensed by the conductive layer) is disposed on one or more of the exposed electrodes of the conductive layer 104 .
  • the sensor chemistry layer 110 is an enzyme layer.
  • the sensor chemistry layer 110 comprises an enzyme capable of producing and/or utilizing oxygen and/or hydrogen peroxide, for example the enzyme glucose oxidase.
  • the enzyme in the sensor chemistry layer is combined with a second carrier protein such as human serum albumin, bovine serum albumin or the like.
  • an enzyme such as glucose oxidase in the sensor chemistry layer 110 reacts with glucose to produce hydrogen peroxide, a compound which then modulates a current at an electrode.
  • this modulation of current depends on the concentration of hydrogen peroxide, and the concentration of hydrogen peroxide correlates to the concentration of glucose, the concentration of glucose can be determined by monitoring this modulation in the current.
  • the hydrogen peroxide is oxidized at a working electrode which is an anode (also termed herein the anodic working electrode), with the resulting current being proportional to the hydrogen peroxide concentration.
  • Such modulations in the current caused by changing hydrogen peroxide concentrations can by monitored by any one of a variety of sensor detector apparatuses such as a universal sensor amperometric biosensor detector or one of the other variety of similar devices known in the art such as glucose monitoring devices produced by Medtronic MiniMed.
  • sensor detector apparatuses such as a universal sensor amperometric biosensor detector or one of the other variety of similar devices known in the art such as glucose monitoring devices produced by Medtronic MiniMed.
  • the analyte sensing layer 110 can be applied over portions of the conductive layer or over the entire region of the conductive layer. Typically the analyte sensing layer 110 is disposed on the working electrode which can be the anode or the cathode. Optionally, the analyte sensing layer 110 is also disposed on a counter and/or reference electrode. While the analyte sensing layer 110 can be up to about 1000 microns ( ⁇ m) in thickness, typically the analyte sensing layer is relatively thin as compared to those found in sensors previously described in the art, and is for example, typically less than 1, 0.5, 0.25 or 0.1 microns in thickness.
  • some methods for generating a thin analyte sensing layer 110 include spin coating processes, dip and dry processes, low shear spraying processes, ink-jet printing processes, silk screen processes and the like. Most typically the thin analyte sensing layer 110 is applied using a spin coating process.
  • the analyte sensing layer 110 is coated with one or more additional layers.
  • the one or more additional layers includes a protein layer 116 disposed upon the analyte sensing layer 110 .
  • the protein layer 116 comprises a protein such as albumin or the like.
  • the protein layer 116 comprises human serum albumin.
  • an additional layer includes an analyte modulating layer 112 that is disposed above the analyte sensing layer 110 to regulate analyte contact with the analyte sensing layer 110 .
  • the analyte modulating membrane layer 112 can comprise a glucose limiting membrane, which regulates the amount of glucose that contacts an enzyme such as glucose oxidase that is present in the analyte sensing layer.
  • glucose limiting membranes can be made from a wide variety of materials known to be suitable for such purposes, e.g., silicone compounds such as polydimethyl siloxanes, polyurethanes, polyurea cellulose acetates, Nafion, polyester sulfonic acids (e.g. Kodak AQ), hydrogels or any other suitable hydrophilic membranes known to those skilled in the art.
  • an adhesion promoter layer 114 is disposed between the analyte modulating layer 112 and the analyte sensing layer 110 as shown in FIG. 2 in order to facilitate their contact and/or adhesion.
  • an adhesion promoter layer 114 is disposed between the analyte modulating layer 112 and the protein layer 116 as shown in FIG. 2 in order to facilitate their contact and/or adhesion.
  • the adhesion promoter layer 114 can be made from any one of a wide variety of materials known in the art to facilitate the bonding between such layers.
  • the adhesion promoter layer 114 comprises a silane compound.
  • protein or like molecules in the analyte sensing layer 110 can be sufficiently crosslinked or otherwise prepared to allow the analyte modulating membrane layer 112 to be disposed in direct contact with the analyte sensing layer 110 in the absence of an adhesion promoter layer 114 .
  • sensors can be designed to contain elements having a combination of some or all of the material properties and/or functions of the elements/constituents discussed below (e.g. an element that serves both as a supporting base constituent and/or a conductive constituent and/or a matrix for the analyte sensing constituent and which further functions as an electrode in the sensor).
  • Sensors of the invention typically include a base constituent (see, e.g. element 102 in FIG. 2 ).
  • base constituent is used herein according to art accepted terminology and refers to the constituent in the apparatus that typically provides a supporting matrix for the plurality of constituents that are stacked on top of one another and comprise the functioning sensor.
  • the base constituent comprises a thin film sheet of insulative (e.g. electrically insulative and/or water impermeable) material.
  • This base constituent can be made of a wide variety of materials having desirable qualities such as water impermeability and hermeticity. Some materials include metallic ceramic and polymeric substrates or the like.
  • the base constituent and/or the conductive constituent can be constructed to produce electrodes having a configuration where the electrochemically reactive surface area of an electrode is disposed on the geometric feature so that the electrochemically reactive surface area is greater than if it was disposed on a flat surface.
  • the base constituent may be self-supporting or further supported by another material as is known in the art.
  • the base constituent 102 comprises a ceramic.
  • the ceramic base comprises a composition that is predominantly Al 2 O 3 (e.g. 96%).
  • the use of alumina as an insulating base constituent for use with implantable devices is disclosed in U.S. Pat. Nos. 4,940,858, 4,678,868 and 6,472,122 which are incorporated herein by reference.
  • the base constituents of the invention can further include other elements known in the art, for example hermetical vias (see, e.g. WO 03/023388).
  • the base constituent can be relatively thick constituent (e.g. thicker than 25 microns).
  • a nonconductive ceramic such as alumina, in thin constituents, e.g., less than about 25 microns.
  • Embodiments of invention disclosed herein provide individual elements and sensors which exhibit a combination of the independent advantages found in each of the two sensor classes disclosed above.
  • a first embodiment of the invention immobilizes an enzyme onto a thick (1-1,000 micron), porous substrate which functions as an electrode in the sensor.
  • the porous electrode is designed to exhibit an increased surface area, for example by constructing it from a lattice of equal-sized adjoining spheres.
  • glucose oxidase is immobilized on a thick (1-1,000 micron), porous metallic substrate that is manufactured from a lattice of equal-sized adjoining spheres and which function as a hydrogen peroxide-consuming electrode.
  • Equation (1) The advantage of such a thick, porous electrode matrices relative to thin, flat electrode matrices is demonstrated using Equation (1):
  • a avail A proj 3 ⁇ L ⁇ ( 1 - ⁇ ) ⁇ ⁇ R ( 1 )
  • the thick, porous electrode is modeled as a lattice of equal-sized adjoining spheres, while the thin electrode is modeled as a two-dimensional surface.
  • the surface area available for enzyme or protein immobilization is A avail
  • the projected area of the electrode is A proj
  • the porosity and thickness of the electrode are L and ⁇ , respectively.
  • the spheres making-up the thick electrode are of radius R, while the fraction of the spheres' surface area available for enzyme or protein immobilization is ⁇ .
  • the porosity range of the such as the porous electrode matrices discussed above is typically 5-99%, 10-99%, 20-99%, 30-99%, 40-99%, 50-99% or 60-99%.
  • the porosity of matrices can be evaluated using any one of a variety of methods known in the art. In certain contexts for example, artisans may wish to examine porosity of a matrix via mercury porosimetry (see, e.g. U.S. Pat. No. 5,609,839), liquid intrusion porosimetry (see, e.g. U.S. Pat. No. 4,660,412), gas porosimetry (see, e.g.
  • nano-porous, micro-porous and macro-porous are used when discussing certain embodiments of the porous matrices that are disclosed herein.
  • platinum-black is commonly used to increase the electrochemically effective surface area of a working electrode.
  • Standard platinum-black electrodes have a great deal of porosity, with the pores being sized so that only very small molecules like H 2 , O 2 , and H 2 O can get inside them.
  • Platinum-black electrodes having this characteristic are termed nano-porous.
  • Such nano-pores have a size range that permits small molecules like H 2 , O 2 , and H 2 O the get inside them, but prevents larger molecules like GOx from getting inside.
  • the electrodes used in the sensors have both nano- and micro-porosity.
  • Micro-pores are characterized in that they are large enough to allow molecules such as GOx to be immobilized inside of them, but are small enough so that any molecule of GOx is relatively close (less than about 0.1, 1, or 2 microns) to the surface of the working electrode. Electrodes having this micro-porosity exhibit a number of desirable characteristics.
  • micro-porous electrodes that allow the placement of immobilized GOx in close proximity to an H 2 O 2 -consuming electrode will increase the lifetime of GOx in the sensor.
  • the hydrogel typically utilized in a variety of analyte sensors is replaced with an essentially rigid, non-swelling porous enzyme-polymer matrix.
  • bio-sensing enzymes can be stably immobilized via covalent bonding to a rigid, macroporous polymer that has optionally been molded into a specified shape.
  • molded continuous rods of macroporous polymers have been developed for use as chromatographic separation media (see, e.g. U.S. Pat. No. 5,453,185 and WO 93/07945).
  • Suitable polymers are essentially incompressible and do not change their overall size in response to changes in their solvating environment.
  • the electrochemical sensors of the invention typically include a conductive constituent disposed upon the base constituent that includes at least one electrode for contacting an analyte or its byproduct (e.g. oxygen and/or hydrogen peroxide) to be assayed (see, e.g. element 104 in FIG. 2 ).
  • a conductive constituent disposed upon the base constituent that includes at least one electrode for contacting an analyte or its byproduct (e.g. oxygen and/or hydrogen peroxide) to be assayed (see, e.g. element 104 in FIG. 2 ).
  • the term “conductive constituent” is used herein according to art accepted terminology and refers to electrically conductive sensor elements such as electrodes which are capable of measuring and a detectable signal and conducting this to a detection apparatus.
  • An illustrative example of this is a conductive constituent that can measure an increase or decrease in current in response to exposure to a stimuli such as the change in the concentration of an analyte or its byproduct as compared to a reference electrode that does not experience the change in the concentration of the analyte, a coreactant (e.g. oxygen) used when the analyte interacts with a composition (e.g. the enzyme glucose oxidase) present in analyte sensing constituent 110 or a reaction product of this interaction (e.g. hydrogen peroxide).
  • a stimuli such as the change in the concentration of an analyte or its byproduct as compared to a reference electrode that does not experience the change in the concentration of the analyte
  • a coreactant e.g. oxygen
  • a composition e.g. the enzyme glucose oxidase
  • a reaction product of this interaction e.g. hydrogen peroxide
  • Illustrative examples of such elements include electrode
  • a working electrode which can be made from non-corroding metal or carbon.
  • a carbon working electrode may be vitreous or graphitic and can be made from a solid or a paste.
  • a metallic working electrode may be made from platinum group metals, including palladium or gold, or a non-corroding metallically conducting oxide, such as ruthenium dioxide.
  • the electrode may comprise a silver/silver chloride electrode composition.
  • the working electrode may be a wire or a thin conducting film applied to a substrate, for example, by coating or printing. Typically, only a portion of the surface of the metallic or carbon conductor is in electrolytic contact with the analyte-containing solution. This portion is called the working surface of the electrode.
  • the remaining surface of the electrode is typically isolated from the solution by an electrically insulating cover constituent 106 .
  • an electrically insulating cover constituent 106 examples include polymers such as polyimides, polytetrafluoroethylene, polyhexafluoropropylene and silicones such as polysiloxanes.
  • the analyte sensors of the invention typically include a reference electrode or a combined reference and counter electrode (also termed a quasi-reference electrode or a counter/reference electrode). If the sensor does not have a counter/reference electrode then it may include a separate counter electrode, which may be made from the same or different materials as 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. These working electrodes in the sensor may be integrally connected or they may be kept separate.
  • the analyte sensors of the present invention are implanted subcutaneously in the skin of a mammal for direct contact with the body fluids of the mammal, such as blood.
  • the sensors can be implanted into other regions within the body of a mammal such as in the intraperotineal space.
  • multiple working electrodes they may be implanted together or at different positions in the body.
  • the counter, reference, and/or counter/reference electrodes may also be implanted either proximate to the working electrode(s) or at other positions within the body of the mammal.
  • the electrochemical sensors of the invention optionally include an interference rejection constituent disposed between the surface of the electrode and the environment to be assayed.
  • an interference rejection constituent disposed between the surface of the electrode and the environment to be assayed.
  • certain sensor embodiments rely on the oxidation and/or reduction of hydrogen peroxide generated by enzymatic reactions on the surface of a working electrode at a constant potential applied. Because amperometric detection based on direct oxidation of hydrogen peroxide requires a relatively high oxidation potential, sensors employing this detection scheme may suffer interference from oxidizable species that are present in biological fluids such as ascorbic acid, uric acid and acetaminophen.
  • interference rejection constituent refers to a coating or membrane in the sensor that functions to inhibit spurious signals generated by such oxidizable species which interfere with the detection of the signal generated by the analyte to be sensed.
  • interference rejection constituents include one or more layers or coatings of compounds such as hydrophilic polyurethanes, cellulose acetate (including cellulose acetate incorporating agents such as poly(ethylene glycol), polyethersulfones, polytetrafluoroethylenes, the perfluoronated ionomer NafionTM, polyphenylenediamine, epoxy and the like.
  • the electrochemical sensors of the invention include an analyte sensing constituent disposed on the electrodes of the sensor (see, e.g. element 110 in FIG. 2 ).
  • analyte sensing constituent is used herein according to art accepted terminology and refers to a constituent comprising a material that is capable of recognizing or reacting with an analyte whose presence is to be detected by the analyte sensor apparatus. Typically this material in the analyte sensing constituent produces a detectable signal after interacting with the analyte to be sensed, typically via the electrodes of the conductive constituent.
  • the analyte sensing constituent and the electrodes of the conductive constituent work in combination to produce the electrical signal that is read by an apparatus associated with the analyte sensor.
  • the analyte sensing constituent comprises an enzyme capable of reacting with and/or producing a molecule whose change in concentration can be measured by measuring the change in the current at an electrode of the conductive constituent (e.g. oxygen and/or hydrogen peroxide), for example the enzyme glucose oxidase.
  • An enzyme capable of producing a molecule such as hydrogen peroxide can be disposed on the electrodes according to a number of processes known in the art.
  • the analyte sensing constituent can coat all or a portion of the various electrodes of the sensor.
  • the analyte sensing constituent may coat the electrodes to an equivalent degree.
  • the analyte sensing constituent may coat different electrodes to different degrees, with for example the coated surface of the working electrode being larger than the coated surface of the counter and/or reference electrode.
  • Typical sensor embodiments of this element of the invention utilize an enzyme (e.g. glucose oxidase) that has been combined with a second protein (e.g. albumin) in a fixed ratio (e.g. one that is typically optimized for glucose oxidase stabilizing properties) and then applied on the surface of an electrode to form a thin enzyme constituent.
  • the analyte sensing constituent comprises a GOx and HSA mixture.
  • the GOx reacts with glucose present in the sensing environment (e.g. the body of a mammal) and generates hydrogen peroxide according to the reaction shown in FIG.
  • a glucose oxidase/albumin is prepared in a physiological solution (e.g., phosphate buffered saline at neutral pH) with the albumin being present in an range of about 0.5%-10% by weight.
  • the stabilized glucose oxidase constituent that is formed on the analyte sensing constituent is very thin as compared to those previously described in the art, for example less than 2, 1, 0.5, 0.25 or 0.1 microns in thickness.
  • One illustrative embodiment of the invention utilizes a stabilized glucose oxidase constituent for coating the surface of an electrode wherein the glucose oxidase is mixed with a carrier protein in a fixed ratio within the constituent, and the glucose oxidase and the carrier protein are distributed in a substantially uniform manner throughout the constituent. Typically the constituent is less than 2 microns in thickness.
  • an enzyme layer can be greater 2 microns.
  • sensors having these extremely thin analyte sensing constituents have material properties that exceed those of sensors having thicker coatings including enhanced longevity, linearity, regularity as well as improved signal to noise ratios. While not being bound by a specific scientific theory, it is believed that sensors having extremely thin analyte sensing constituents have surprisingly enhanced characteristics as compared to those of thicker constituents because in thicker enzyme constituents only a fraction of the reactive enzyme within the constituent is able to access the analyte to be sensed. In sensors utilizing glucose oxidase, the thick coatings produced by electrodeposition may hinder the ability of hydrogen peroxide generated at the reactive interface of a thick enzyme constituent to contact the sensor surface and thereby generate a signal.
  • the enzyme and the second protein are typically treated to form a crosslinked matrix (e.g. by adding a cross-linking agent to the protein mixture).
  • crosslinking conditions may be manipulated to modulate factors such as the retained biological activity of the enzyme, its mechanical and/or operational stability.
  • Illustrative crosslinking procedures are described in U.S. patent application Ser. No. 10/335,506 and PCT publication WO 03/035891 which are incorporated herein by reference.
  • an amine cross-linking reagent such as, but not limited to, glutaraldehyde, can be added to the protein mixture.
  • glutaraldehyde can be added to the protein mixture.
  • the addition of a cross-linking reagent to the protein mixture creates a protein paste.
  • the concentration of the cross-linking reagent to be added may vary according to the concentration of the protein mixture. While glutaraldehyde is an illustrative crosslinking reagent, other cross-linking reagents may also be used or may be used in place of glutaraldehyde, including, but not limited to, an amine reactive, homofunctional, cross-linking reagent such as Disuccinimidyl Suberate (DSS). Another example is 1-Ethyl-3 (3-Dimethylaminopropyl) Carbodiimide (EDC), which is a zero-length cross-linker. EDC forms an amide bond between carboxylic acid and amine groups. Other suitable cross-linkers also may be used, as will be evident to those skilled in the art.
  • DSS Disuccinimidyl Suberate
  • EDC 1-Ethyl-3 (3-Dimethylaminopropyl) Carbodiimide
  • EDC forms an amide bond between carboxylic acid and amine groups.
  • the GOx and/or carrier protein concentration may vary for different embodiments of the invention.
  • the GOx concentration may be within the range of approximately 50 mg/ml (approximately 10,000 U/ml) to approximately 700 mg/ml (approximately 150,000 U/ml).
  • the GOx concentration is about 115 mg/ml (approximately 22,000 U/ml).
  • the HSA concentration may vary between about 0.5%-30% (w/v), depending on the GOx concentration.
  • the HSA concentration is about 1-10% w/v, and most typically is about 5% w/v.
  • collagen or BSA or other structural proteins used in these contexts can be used instead of or in addition to HSA.
  • GOx is discussed as an illustrative enzyme in the analyte sensing constituent, other proteins and/or enzymes may also be used or may be used in place of GOx, including, but not limited to glucose dehydrogenase or hexokinase, hexose oxidase, lactate oxidase, and the like. Other proteins and/or enzymes may also be used, as will be evident to those skilled in the art.
  • HSA is employed in the example embodiment, other structural proteins, such as BSA, collagens or the like, could be used instead of or in addition to HSA.
  • concentrations other than those discussed herein may be utilized. For example, depending on the enzyme employed, concentrations ranging from approximately 10% weight per weight to 70% weight per weight may be suitable. The concentration may be varied not only depending on the particular enzyme being employed, but also depending on the desired properties of the resulting protein matrix. For example, a certain concentration may be utilized if the protein matrix is to be used in a diagnostic capacity while a different concentration may be utilized if certain structural properties are desired. Those skilled in the art will understand that the concentration utilized may be varied through experimentation to determine which concentration (and of which enzyme or protein) may yield the desired result.
  • the analyte sensing constituent includes a composition (e.g. glucose oxidase) capable of producing a signal (e.g. a change in oxygen and/or hydrogen peroxide concentrations) that can be sensed by the electrically conductive elements (e.g. electrodes which sense changes in oxygen and/or hydrogen peroxide concentrations).
  • a composition e.g. glucose oxidase
  • the electrically conductive elements e.g. electrodes which sense changes in oxygen and/or hydrogen peroxide concentrations
  • the composition comprises an enzyme that modulates hydrogen peroxide concentrations upon reaction with an analyte to be sensed.
  • the composition comprises an enzyme that modulates oxygen concentrations upon reaction with an analyte to be sensed.
  • an enzyme that modulates oxygen concentrations upon reaction with an analyte to be sensed.
  • a wide variety of enzymes that either use or produce hydrogen peroxide and/or oxygen in a reaction with a physiological analyte are known in the art and these enzymes can be readily incorporated into the analyte sensing constituent composition.
  • a variety of other enzymes known in the art can produce and/or utilize compounds whose modulation can be detected by electrically conductive elements such as the electrodes that are incorporated into the sensor designs described herein.
  • Such enzymes include for example, enzymes specifically described in Table 1, pages 15-29 and/or Table 18, pages 111-112 of Protein Immobilization: Fundamentals and Applications (Bioprocess Technology, Vol 14) by Richard F. Taylor (Editor) Publisher: Marcel Dekker; (Jan. 7, 1991) the entire contents of which are incorporated herein by reference.
  • analyte sensing constituents can be formed to include antibodies whose interaction with a target analyte is capable of producing a detectable signal that can be sensed by the electrically conductive elements after interacting with the target analyte whose presence is to be detected.
  • U.S. Pat. No. 5,427,912 (which is incorporated herein by reference) describes an antibody-based apparatus for electrochemically determining the concentration of an analyte in a sample.
  • a mixture which includes the sample to be tested, an enzyme-acceptor polypeptide, an enzyme-donor polypeptide linked to an analyte analog (enzyme-donor polypeptide conjugate), a labeled substrate, and an antibody specific for the analyte to be measured.
  • the analyte and the enzyme-donor polypeptide conjugate competitively bind to the antibody.
  • the enzyme-donor polypeptide conjugate is not bound to antibody, it will spontaneously combine with the enzyme acceptor polypeptide to form an active enzyme complex.
  • the active enzyme then hydrolyzes the labeled substrate, resulting in the generation of an electroactive label, which can then be oxidized at the surface of an electrode.
  • a current resulting from the oxidation of the electroactive compound can be measured and correlated to the concentration of the analyte in the sample.
  • U.S. Pat. No. 5,149,630 (which is incorporated herein by reference) describes an electrochemical specific binding assay of a ligand (e.g., antigen, hapten or antibody) wherein at least one of the components is enzyme-labelled, and which includes the step of determining the extent to which the transfer of electrons between the enzyme substrate and an electrode, associated with the substrate reaction, is perturbed by complex formation or by displacement of any ligand complex relative to unbound enzyme-labelled component. The electron transfer is aided by electron-transfer mediators which can accept electrons from the enzyme and donate them to the electrode or vice versa (e.g.
  • Anti-LDH5 antibody is bound to a suitable glassy carbon electrode; this is contacted with the substrate containing LDH5, rinsed, inserted into a NAD solution, connected to an amperometric system, current changes are measured in the presence of differing concentrations of lactic acid, which are indicative of the quantity of LDH-5.
  • U.S. Pat. No. 6,410,251 (which is incorporated herein by reference) describes an apparatus and method for detecting or assaying one constituting member in a specific binding pair; for example, the antigen in an antigen/antibody pair, by utilizing specific binding such as binding between an antigen and an antibody, together with redox reaction for detecting a label, wherein an oxygen micro-electrode with a sensing surface area is used.
  • U.S. Pat. No. 4,402,819 (which is incorporated herein by reference) describes an antibody-selective potentiometric electrode for the quantitative determination of antibodies (as the analyte) in dilute liquid serum samples employing an insoluble membrane incorporating an antigen having bonded thereto an ion carrier effecting the permeability of preselected cations therein, which permeability is a function of specific antibody concentrations in analysis, and the corresponding method of analysis.
  • analyte sensing constituents of the sensors disclosed herein include polymers that bind specific types of cells or cell components (e.g. polypeptides, carbohydrates and the like); single-strand DNA; antigens and the like.
  • the detectable signal can be, for example, an optically detectable change, such as a color change or a visible accumulation of the desired analyte (e.g., cells).
  • Sensing elements can also be formed from materials that are essentially non-reactive (i.e., controls).
  • sensor elements are beneficially included, for example, in sensors for use in cell-sorting assays and assays for the presence of pathogenic organisms, such as viruses (HIV, hepatitis-C, etc.), bacteria, protozoa and the like.
  • pathogenic organisms such as viruses (HIV, hepatitis-C, etc.), bacteria, protozoa and the like.
  • analyte sensors that measure an analyte that is present in the external environment and that can in itself produce a measurable change in current at an electrode.
  • the analyte sensing constituent can be optional.
  • the electrochemical sensors of the invention optionally include a protein constituent disposed between the analyte sensing constituent and the analyte modulating constituent (see, e.g. element 116 in FIG. 2 ).
  • the term “protein constituent” is used herein according to art accepted terminology and refers to constituent containing a carrier protein or the like that is selected for compatibility with the analyte sensing constituent and/or the analyte modulating constituent.
  • the protein constituent comprises an albumin such as human serum albumin.
  • the HSA concentration may vary between about 0.5%-30% (w/v). Typically the HSA concentration is about 1-10% w/v, and most typically is about 5% w/v.
  • collagen or BSA or other structural proteins used in these contexts can be used instead of or in addition to HSA.
  • This constituent is typically crosslinked on the analyte sensing constituent according to art accepted protocols.
  • the electrochemical sensors of the invention can include one or more adhesion promoting (AP) constituents (see, e.g. element 114 in FIG. 2 ).
  • AP adhesion promoting
  • the term “adhesion promoting constituent” is used herein according to art accepted terminology and refers to a constituent that includes materials selected for their ability to promote adhesion between adjoining constituents in the sensor.
  • the adhesion promoting constituent is disposed between the analyte sensing constituent and the analyte modulating constituent.
  • the adhesion promoting constituent is disposed between the optional protein constituent and the analyte modulating constituent.
  • the adhesion promoter constituent can be made from any one of a wide variety of materials known in the art to facilitate the bonding between such constituents and can be applied by any one of a wide variety of methods known in the art.
  • the adhesion promoter constituent comprises a silane compound such as ⁇ -aminopropyltrimethoxysilane.
  • silane coupling reagents especially those of the formula R′Si(OR) 3 in which R′ is typically an aliphatic group with a terminal amine and R is a lower alkyl group
  • R′ is typically an aliphatic group with a terminal amine and R is a lower alkyl group
  • R′ is typically an aliphatic group with a terminal amine and R is a lower alkyl group
  • chemically modified electrodes in which a silane such as ⁇ -aminopropyltriethoxysilane and glutaraldehyde were used in a step-wise process to attach and to co-crosslink bovine serum albumin (BSA) and glucose oxidase (GO X ) to the electrode surface are well known in the art (see, e.g. Yao, T. Analytica Chim. Acta 1983, 148, 27-33).
  • BSA bovine serum albumin
  • GO X glucose oxidase
  • the adhesion promoting constituent further comprises one or more compounds that can also be present in an adjacent constituent such as the polydimethyl siloxane (PDMS) compounds that serves to limit the diffusion of analytes such as glucose through the analyte modulating constituent.
  • the formulation comprises 0.5-20% PDMS, typically 5-15% PDMS, and most typically 10% PDMS.
  • the adhesion promoting constituent includes an agent selected for its ability to crosslink a siloxane moiety present in a proximal constituent such as the analyte modulating constituent.
  • the adhesion promoting constituent includes an agent selected for its ability to crosslink an amine or carboxyl moiety of a protein present in a proximal constituent such a the analyte sensing constituent and/or the protein constituent.
  • the electrochemical sensors of the invention include an analyte modulating constituent disposed on the sensor (see, e.g. element 112 in FIG. 2 ).
  • analyte modulating constituent is used herein according to art accepted terminology and refers to a constituent that typically forms a membrane on the sensor that operates to modulate the diffusion of one or more analytes, such as glucose, through the constituent.
  • the analyte modulating constituent is an analyte-limiting membrane which operates to prevent or restrict the diffusion of one or more analytes, such as glucose, through the constituents.
  • the analyte-modulating constituent operates to facilitate the diffusion of one or more analytes, through the constituents.
  • such analyte modulating constituents can be formed to prevent or restrict the diffusion of one type of molecule through the constituent (e.g. glucose), while at the same time allowing or even facilitating the diffusion of other types of molecules through the constituent (e.g. O 2 ).
  • one type of molecule e.g. glucose
  • other types of molecules e.g. O 2
  • glucose and oxygen from blood diffuse through a primary membrane of the sensor.
  • an enzyme such as glucose oxidase, catalyzes the conversion of glucose to hydrogen peroxide and gluconolactone.
  • the hydrogen peroxide may diffuse back through the analyte modulating constituent, or it may diffuse to an electrode where it can be reacted to form oxygen and a proton to produce a current that is proportional to the glucose concentration.
  • the sensor membrane assembly serves several functions, including selectively allowing the passage of glucose therethrough.
  • an illustrative analyte modulating constituent is a semi-permeable membrane which permits passage of water, oxygen and at least one selective analyte and which has the ability to absorb water, the membrane having a water soluble, hydrophilic polymer.
  • analyte modulating composition includes PDMS.
  • the analyte modulating constituent includes an agent selected for its ability to crosslink a siloxane moiety present in a proximal constituent.
  • the adhesion promoting constituent includes an agent selected for its ability to crosslink an amine or carboxyl moiety of a protein present in a proximal constituent.
  • the electrochemical sensors of the invention include one or more cover constituents which are typically electrically insulating protective constituents (see, e.g. element 106 in FIG. 2 ). Typically, such cover constituents are disposed on at least a portion of the analyte modulating constituent.
  • Acceptable polymer coatings for use as the insulating protective cover constituent can include, but are not limited to, non-toxic biocompatible polymers such as silicone compounds, polyimides, biocompatible solder masks, epoxy acrylate copolymers, or the like. Further, these coatings can be photo-imageable to facilitate photolithographic forming of apertures through to the conductive constituent.
  • a typical cover constituent comprises spun on silicone. As is known in the art, this constituent can be a commercially available RTV (room temperature vulcanized) silicone composition. A typical chemistry in this context is polydimethyl siloxane (acetoxy based).
  • the analyte sensor apparatus disclosed herein has a number of embodiments.
  • a general embodiment of the invention is an analyte sensor apparatus for implantation within a mammal. While the analyte sensors are typically designed to be implantable within the body of a mammal, the sensors are not limited to any particular environment can instead be used in a wide variety of contexts, for example for the analysis of most liquid samples including biological fluids such as whole-blood, lymph, plasma, serum, saliva, urine, stool, perspiration, mucus, teats, cerebrospinal fluid, nasal secretion, cervical or vaginal secretion, semen, pleural fluid, amniotic fluid, peritoneal fluid, middle ear fluid, joint fluid, gastric aspirate or the like.
  • biological fluids such as whole-blood, lymph, plasma, serum, saliva, urine, stool, perspiration, mucus, teats, cerebrospinal fluid, nasal secretion, cervical or vaginal secretion, semen,
  • the sensor embodiments disclosed herein can be used to sense analytes of interest in one or more physiological environments.
  • the sensor can be in direct contact with interstitial fluids as typically occurs with subcutaneous sensors.
  • the sensors of the present invention may also be part of a skin surface system where interstitial glucose is extracted through the skin and brought into contact with the sensor (see, e.g. U.S. Pat. Nos. 6,155,992 and 6,706,159 which are incorporated herein by reference).
  • the sensor can be in contact with blood as typically occurs for example with intravenous sensors.
  • the sensor embodiments of the invention further include those adapted for use in a variety of contexts.
  • the senor can be designed for use in mobile contexts, such as those employed by ambulatory users.
  • the sensor can be designed for use in stationary contexts such as those adapted for use in clinical settings.
  • Such sensor embodiments include, for example, those used to monitor one or more analytes present in one or more physiological environments in a hospitalized patient.
  • Sensors of the invention can also be incorporated in to a wide variety of medical systems known in the art. Sensors of the invention can be used, for example, in a closed loop infusion systems designed to control the rate that medication is infused into the body of a user.
  • a closed loop infusion system can include a sensor and an associated meter which generates an input to a controller which in turn operates a delivery system (e.g. one that calculates a dose to be delivered by a medication infusion pump).
  • the meter associated with the sensor may also transmit commands to, and be used to remotely control, the delivery system.
  • the senor is a subcutaneous sensor in contact with interstitial fluid to monitor the glucose concentration in the body of the user, and the liquid infused by the delivery system into the body of the user includes insulin.
  • Illustrative systems are disclosed for example in U.S. Pat. Nos. 6,558,351 and 6,551,276; PCT Application Nos. US99/21703 and US99/22993; as well as WO 2004/008956 and WO 2004/009161, all of which are incorporated herein by reference.
  • Certain embodiments of the invention measure peroxide and have the advantageous characteristic of being suited for implantation in a variety of sites in the mammal including regions of subcutaneous implantation and intravenous implantation as well as implantation into a variety of non-vascular regions.
  • a peroxide sensor design that allows implantation into non-vascular regions has advantages over certain sensor apparatus designs that measure oxygen due to the problems with oxygen noise that can occur in oxygen sensors implanted into non-vascular regions. For example, in such implanted oxygen sensor apparatus designs, oxygen noise at the reference sensor can compromise the signal to noise ratio which consequently perturbs their ability to obtain stable glucose readings in this environment.
  • the peroxide sensors of the invention therefore overcome the difficulties observed with such oxygen sensors in non-vascular regions.
  • Certain peroxide sensor embodiments of the invention further include advantageous long term or “permanent” sensors which are suitable for implantation in a mammal for a time period of greater than 30 days.
  • medical devices such as the sensors described herein can be categorized into three groups based on implant duration: (1) “Limited” ( ⁇ 24 hours), (2) “Prolonged” (24 hours-30 days), and (3) “Permanent” (>30 days).
  • the design of the peroxide sensor of the invention allows for a “Permanent” implantation according to this categorization, i.e. >30 days.
  • the highly stable design of the peroxide sensor of the invention allows for an implanted sensor to continue to function in this regard for 2, 3, 4, 5, 6 or 12 or more months.
  • the analyte sensor apparatus structure comprises a base layer and a conductive layer disposed upon the base layer (e.g. a porous matrix) and functions as one or more electrodes.
  • the conductive layer can include a working electrode, a reference electrode and/or a counter electrode. These electrodes can be spaced in proximity, or alternatively are spaced distally, according to the specific design.
  • the sensor apparatus design is such that certain electrodes (e.g. the working electrode) can be exposed to the solution containing the analyte to be sensed (e.g. via an aperture) in the sensor apparatus.
  • the sensor apparatus design is such that certain electrodes (e.g. the reference electrode) are not exposed to the solution containing the analyte to be sensed in the sensor apparatus.
  • compositions for use in biosensors.
  • Such compositions are typically designed to implantable within a mammal and comprise a porous matrix having a surface coated with an immobilized enzyme, for example glucose oxidase, glucose dehydrogenase, lactate oxidase, hexokinase or lactate dehydrogenase.
  • an immobilized enzyme for example glucose oxidase, glucose dehydrogenase, lactate oxidase, hexokinase or lactate dehydrogenase.
  • the porous matrix coated with an immobilized enzyme is capable of acting as an electrode in an electrochemical sensor.
  • the electrode in the electrochemical sensor consumes hydrogen peroxide.
  • porous matrices used in various embodiments of the biosensors of the invention can be generated from a variety of materials and can be adapted to a variety of compositional configurations.
  • the porous matrix comprises a ceramic material and/or a metal and/or a macroporous polymer.
  • the porous matrix comprises a lattice of particles. Typically the particles are spherical.
  • porous matrix has a surface area that is at least 2, 4, 6, 8, 10, 12, 14, 16 or 18 times the surface area of a non-porous matrix of same dimensions.
  • the porous matrix is at least 1, 10, 100, or 1000 microns thick.
  • the porosity range of the porous matrix is optionally about 5-99.9% and typically is about 40-99%.
  • the porosity of these matrices can be measured by one of the protocols typically used in the art such as mercury or gas porosimetry, size-exclusion chromatography using marker molecules of various sizes and molecular weights (e.g. acetone, various globular proteins of a defined size, blue dextran), and cyclic voltammetry.
  • a related embodiment of the invention is an analyte sensor apparatus for implantation within a mammal which includes a porous matrix having a surface coated with an immobilized enzyme, for example glucose oxidase.
  • the porous matrix comprises a working electrode; and the immobilized enzyme is disposed within an analyte sensing layer disposed on the working electrode, such that the analyte sensing layer detectably alters the electrical current at the working electrode in the conductive layer in the presence of an analyte.
  • the sensor further comprises an analyte modulating layer disposed on the analyte sensing layer, wherein the analyte modulating layer modulates the diffusion of the analyte therethrough.
  • the senor further comprises an adhesion promoting layer disposed on the analyte sensing layer, wherein the adhesion promoting layer promotes the adhesion between the analyte sensing layer and an analyte modulating layer disposed on the analyte sensing layer.
  • the sensor further comprises a protein layer disposed between the analyte sensing layer and the analyte modulating layer.
  • the sensor further comprises a cover layer disposed on at least a portion of the analyte modulating layer, wherein the cover layer further includes an aperture that exposes at least a portion of the analyte modulating layer to a solution comprising the analyte to be sensed.
  • a related embodiment of the invention is a method of making a sensor apparatus for implantation within a mammal comprising the steps of providing a layer comprising a porous matrix, forming an analyte sensing layer on the porous matrix, wherein the analyte sensing layer includes an enzyme such as glucose oxidase that can alter the electrical current at the surface of the porous matrix in the presence of an analyte so that the porous matrix having the analyte sensing layer formed thereon functions as an electrode.
  • an enzyme such as glucose oxidase
  • Such methods further include the steps of optionally forming a protein layer on the analyte sensing layer, forming an adhesion promoting layer on the analyte sensing layer or the optional protein layer, forming an analyte modulating layer disposed on the adhesion promoting layer, wherein the analyte modulating layer includes a composition that modulates the diffusion of the analyte therethrough; and forming a cover layer disposed on at least a portion of the analyte modulating layer, wherein the cover layer further includes an aperture over at least a portion of the analyte modulating layer.
  • Another embodiment of the invention is a method of sensing an analyte within the body of a mammal, the method comprising implanting an analyte sensor into the mammal, the analyte sensor comprising a porous matrix having an analyte sensing layer disposed thereon, wherein the analyte sensing layer detectably alters the electrical current at the surface of the porous matrix in the presence of an analyte so that the porous matrix having the analyte sensing layer formed thereon functions as an electrode, an optional protein layer disposed on the analyte sensing layer, an adhesion promoting layer disposed on the analyte sensing layer or the optional protein layer, wherein the adhesion promoting layer promotes the adhesion between the analyte sensing layer and an analyte modulating layer disposed on the analyte sensing layer, and an analyte modulating layer disposed on the analyte sensing layer, wherein the ana
  • Yet another embodiment of the invention is a method of immobilizing a protein on a rigid macroporous polymer comprising the steps of: combining the protein with the rigid macroporous polymer having functional moieties capable of crosslinking to a protein; and then adding a crosslinking agent capable of immobilizing the protein on the rigid macroporous polymer by crosslinking the functional moieties of the protein with the functional moieties of the rigid macroporous polymer so that the protein is immobilized on the rigid macroporous polymer.
  • the rigid macroporous polymer having functional moieties capable of crosslinking to a protein is made by combining a rigid macroporous polymer having reactive epoxide moieties with a nucleophilic compound so that a rigid macroporous polymer having functional moieties capable of crosslinking to a protein is made.
  • Yet another embodiment of the invention is a method of immobilizing a protein on a rigid macroporous polymer comprising combining a protein having a sulfhydryl, amine, carboxyl or hydroxyl moiety with a rigid macroporous polymer having reactive epoxide moieties under reaction conditions that allow a nucleophilic reaction to occur between the sulfhydryl, amine, carboxyl or hydroxyl moieties on the protein and the epoxide moieties on the rigid macroporous polymer so that the protein is immobilized on the rigid macroporous polymer.
  • at least one nucleophilic moiety on the protein is blocked prior to combining the protein with the rigid macroporous polymer.
  • Analyte sensors of the invention typically incorporate the porous matrices disclosed herein.
  • the analyte sensor apparatus includes an analyte sensing layer disposed on a conductive layer of the sensor, typically covering a portion or all of the working electrode.
  • This analyte sensing layer detectably alters the electrical current at the working electrode in the conductive layer in the presence of an analyte to be sensed.
  • this analyte sensing layer typically includes an enzyme or antibody molecule or the like that reacts with the analyte of interest in a manner that changes the concentrations of a molecule that can modulate the current at the working electrode (see e.g. oxygen and/or hydrogen peroxide as shown in the reaction scheme of FIG.
  • Illustrative analyte sensing layers comprise an enzyme such as glucose oxidase (e.g. for use in glucose sensors) or lactate oxidase (e.g. for use in lactate sensors).
  • the analyte sensing layer is disposed upon a porous metallic and/or ceramic and/or polymeric matrix with this combination of elements functioning as an electrode in the sensor.
  • the analyte-sensing layer further comprises a carrier protein in a substantially fixed ratio with the analyte sensing compound (e.g. the enzyme) and the analyte sensing compound and the carrier protein are distributed in a substantially uniform manner throughout the analyte sensing layer.
  • the analyte sensing layer is very thin, for example, less than 1, 0.5, 0.25 or 0.1 microns in thickness.
  • sensors having such thin analyte sensing layers have surprisingly enhanced characteristics as compared to the thicker layers that are typically generated by electrodeposition because electrodeposition produces 3-5 micron thick enzyme layers in which only a fraction of the reactive enzyme within the coating layer is able to access the analyte to be sensed.
  • Such thicker glucose oxidase pellets that are produced by electrodeposition protocols are further observed to have a poor mechanical stability (e.g. a tendency to crack) and further take a longer time to prepare for actual use, typically taking weeks of testing before it is ready for implantation.
  • these problems are not observed with the thin layered enzyme coatings described herein, these thin coatings are typical embodiments of the invention.
  • the thick coatings produced by electrodeposition may hinder the ability of hydrogen peroxide generated at the reactive interface of the 3-5 micron thick enzyme layer to contact the sensor surface and thereby generate a signal.
  • hydrogen peroxide that is unable to reach a sensor surface due to such thick coatings can diffuse away from the sensor into the environment in which the sensor is placed, thereby decreasing the sensitivity and/or biocompatibility of such sensors.
  • sensors having such thin analyte sensing layers have unexpectedly advantageous properties that result from the fact that processes such as spin coating, or the like, allow for a precise control over the enzyme coating's ratio of glucose oxidase to albumin (which is used as a carrier protein to stabilize the glucose oxidase in the enzyme layer).
  • processes such as spin coating, or the like
  • electrodeposition processes may result in a surface coating in which an optimally determined ratio of enzyme to carrier protein is detrimentally altered in the electrodeposition process, and further wherein the glucose oxidase and the carrier protein are not distributed in a substantially uniform manner throughout the disposed enzyme layer.
  • sensors having such thin analyte sensing layers have unexpectedly faster response times. While not being bound by a specific scientific theory, it is believed that these surprising and advantageous properties result from the observation that thin enzyme layers allow better access to the working electrode surface and may allow a greater proportion of the molecules that modulate current at the electrode to access the electrode surface.
  • an alteration in current in response to exposure to the analyte present in the body of the mammal can be detected via an amperometer within 15, 10, 5 or 2 minutes of the analyte contacting the analyte sensor.
  • the analyte sensing layer has a protein layer disposed thereon and which is typically between this analyte sensing layer and the analyte modulating layer.
  • a protein within the protein layer is an albumin selected from the group consisting of bovine serum albumin and human serum albumin. Typically this protein is crosslinked. Without being bound by a specific scientific theory, it is believed that this separate protein layer enhances sensor function and provides surprising functional benefits by acting as a sort of capacitor that diminishes sensor noise (e.g. spurious background signals).
  • some amount of moisture may form under the analyte modulating membrane layer of the sensor, the layer which regulates the amount of analyte that can contact the enzyme of the analyte sensing layer.
  • This moisture may create a compressible layer that shifts within the sensor as a patient using the sensor moves.
  • Such shifting of layers within the sensor may alter the way that an analyte such as glucose moves through the analyte sensing layers in a manner that is independent of actual physiological analyte concentrations, thereby generating noise.
  • the protein layer may act as a capacitor by protecting an enzyme such as GOx from contacting the moisture layer.
  • This protein layer may confer a number of additional advantages such as promoting the adhesion between the analyte sensing layer and the analyte modulating membrane layer.
  • the presence of this layer may result in a greater diffusion path for molecules such as hydrogen peroxide, thereby localizing it to the electrode sensing element and contributing to an enhanced sensor sensitivity.
  • the analyte sensing layer and/or the protein layer disposed on the analyte sensing layer has an adhesion promoting layer disposed thereon.
  • adhesion promoting layers promote the adhesion between the analyte sensing layer and a proximal layer, typically an analyte modulating layer.
  • This adhesion promoting layer typically comprises a silane compound such as ⁇ -aminopropyltrimethoxysilane which is selected for its ability to promote optimized adhesion between the various sensor layers and functions to stabilize the sensor.
  • sensors having such a silane containing adhesion promoting layers exhibit unexpected properties including an enhanced overall stability.
  • silane containing adhesion promoting layers provide a number of advantageous characteristics in addition to an ability to enhancing sensor stability, and can, for example, play a beneficial role in interference rejection as well as in controlling the mass transfer of one or more desired analytes.
  • the adhesion promoting layer further comprises one or more compounds that can also be present in an adjacent layer such as the polydimethyl siloxane (PDMS) compounds that serves to limit the diffusion of analytes such as glucose through the analyte modulating layer.
  • PDMS polydimethyl siloxane
  • the addition of PDMS to the AP layer for example can be advantageous in contexts where it diminishes the possibility of holes or gaps occurring in the AP layer as the sensor is manufactured.
  • the adhesion promoting layer has an analyte modulating layer disposed thereon which functions to modulate the diffusion of analytes therethrough.
  • the analyte modulating layer includes compositions (e.g. polymers and the like) which serve to enhance the diffusion of analytes (e.g. oxygen) through the sensor layers and consequently function to enrich analyte concentrations in the analyte sensing layer.
  • the analyte modulating layer includes compositions which serve to limit the diffusion of analytes (e.g. glucose) through the sensor layers and consequently function to limit analyte concentrations in the analyte sensing layer.
  • analyte modulating layer includes compositions which serve to limit the diffusion of analytes (e.g. glucose) through the sensor layers and consequently function to limit analyte concentrations in the analyte sensing layer.
  • a hydrophilic glucose limiting membrane i.e. functions to limit the diffusion of glucose therethrough
  • the analyte modulating layer further comprises one or more cover layers which are typically electrically insulating protective layers disposed on at least a portion of the sensor apparatus (e.g. covering the analyte modulating layer).
  • Acceptable polymer coatings for use as the insulating protective cover layer can include, but are not limited to, non-toxic biocompatible polymers such as silicone compounds, polyimides, biocompatible solder masks, epoxy acrylate copolymers, or the like.
  • An illustrative cover layer comprises spun on silicone.
  • the cover layer further includes an aperture that exposes at least a portion of a sensor layer (e.g. analyte modulating layer) to a solution comprising the analyte to be sensed.
  • the analyte sensors described herein can be polarized cathodically to detect, for example, changes in current at the working cathode that result from the changes in oxygen concentration proximal to the working cathode that occur as glucose interacts with glucose oxidase as shown in FIG. 1 .
  • the analyte sensors described herein can be polarized anodically to detect for example, changes in current at the working anode that result from the changes in hydrogen peroxide concentration proximal to the working anode that occur as glucose interacts with glucose oxidase as shown in FIG. 1 .
  • the current at the working electrode(s) is compared to the current at a reference electrode(s) (a control), with the differences between these measurements providing a value that can then be correlated to the concentration of the analyte being measured.
  • Analyte sensor designs that obtain a current value by obtaining a measurement from a comparison of the currents at these dual electrodes are commonly termed, for example, dual oxygen sensors.
  • the analyte sensor apparatus is designed to function via anodic polarization such that the alteration in current is detected at the anodic working electrode in the conductive layer of the analyte sensor apparatus.
  • Structural design features than can be associated with anodic polarization include designing an appropriate sensor configuration comprising a working electrode which is an anode, a counter electrode which is a cathode and a reference electrode and then selectively disposing the appropriate analyte sensing layer on the appropriate portion of the surface of the anode within this design configuration.
  • this anodic polarization structural design includes anodes, cathodes and/or working electrodes having different sized surface areas.
  • this structural design includes features where the working electrode (anode) and/or the coated surface of the working electrode is larger than the counter electrode (cathode) and/or the coated surface of the counter electrode.
  • the alteration in current that can be detected at the anodic working electrode is then correlated with the concentration of the analyte.
  • the working electrode is measuring and utilizing hydrogen peroxide in the oxidation reaction (see e.g. FIG. 1 ), hydrogen peroxide that is produced by an enzyme such as glucose oxidase or lactate oxidase upon reaction with glucose or lactate respectively.
  • Such embodiments of the invention relating to electrochemical glucose and/or lactate sensors having such hydrogen peroxide recycling capabilities are particularly interesting because the recycling of this molecule reduces the amount of hydrogen peroxide that can escape from the sensor into the environment in which it is placed.
  • implantable sensors that are designed to reduce the release of tissue irritants such as hydrogen peroxide will have improved biocompatibility profiles.
  • hydrogen peroxide can react with enzymes such as glucose oxidase and compromise their biological function, such sensors are desired due to their avoidance of this phenomena.
  • the analyte modulating layer e.g. a glucose limiting layer
  • analyte sensors of the invention that comprise a base layer, a conductive layer, an analyte sensing layer, an optional protein layer, an adhesion promoting layer, an analyte modulating layer and a cover layer exhibit a number of unexpected properties.
  • sensors that are structured to function via anodic polarization versus those structured to function via cathodic polarization differences in the electrochemical reactions in the analyte sensing layer as well as at the electrode surface generate and/or consume different chemical entities, thereby altering the chemical environment in which the various sensor elements function in different polarities.
  • the sensor structure disclosed herein provides a surprisingly versatile device that is shown to function with an unexpected degree of stability under a variety of different chemical and/or electrochemical conditions.
  • the sensor layer has a plurality of electrodes including a working electrode (e.g. an anode) and a counter electrode (e.g. a cathode), both of which are coated with an analyte sensing layer comprising an enzyme such as glucose oxidase or lactate oxidase.
  • a working electrode e.g. an anode
  • a counter electrode e.g. a cathode
  • an enzyme such as glucose oxidase or lactate oxidase.
  • Such sensor designs have surprising properties including an enhanced sensitivity. Without being bound by a specific theory, these properties may result from the enhanced oxidation of hydrogen peroxide at the surface of a working or a counter electrode which produces additional oxygen that can be utilized in the glucose sensing reaction (see, e.g., FIG. 1 ).
  • this recycling effect may reduce the oxygen dependent limitations of certain sensor embodiments disclosed herein. Moreover, this design may result in a sensor having a working electrode that can readily reduce available hydrogen peroxide and consequently have a lower electrode potential. Sensors designed to function with lower electrode potentials are typical embodiments of the invention because high electrode potentials in sensors of this type can result in a gas producing hydrolysis reaction which can destabilize the sensors (due to the disruption of sensor layers from gas bubbles produced by hydrolysis reactions).
  • the hydrogen peroxide generated in the enzymatic reaction is very close to the reactive surface of the counter electrode. This can increase the overall efficiency of the sensor in a manner that allows for the production of compact sensor designs which include for example, counter electrodes with smaller reactive surfaces.
  • a specific illustrative example of an analyte sensor apparatus for implantation within a mammal is a peroxide sensor of the following design.
  • a first layer of the peroxide sensor apparatus is a base layer, typically made from a ceramic such as alumina.
  • a subsequent layer disposed upon the base layer is a conductive layer including a plurality of electrodes including an anodic working electrode and a reference electrode.
  • a subsequent layer disposed on the conductive layer is an analyte sensing layer that includes crosslinked glucose oxidase which senses glucose and consequently generates hydrogen peroxide as shown in FIG. 1 .
  • the anodic working electrode experiences a measurable increase in current as the hydrogen peroxide generated contacts this anode in the conductive layer and is oxidized.
  • the reference electrode serves as a control and is physically isolated from the working electrode and the hydrogen peroxide generated according to the reaction shown in FIG. 1 .
  • This analyte sensing layer is typically less than 1, 0.5, 0.25 or 0.1 microns in thickness and comprises a mixture of crosslinked human serum albumin in a substantially fixed ratio with the crosslinked glucose oxidase, with the glucose oxidase and the human serum albumin being distributed in a substantially uniform manner throughout the sensor layer.
  • a subsequent layer disposed on the sensor layer is a protein layer comprising crosslinked human serum albumin.
  • a subsequent layer disposed on the protein layer is an adhesion promoting layer which promotes the adhesion between the analyte sensing layer and/or the protein layer and an analyte modulating layer which is disposed upon these layers.
  • This adhesion promoting layer comprises a silane composition.
  • a subsequent layer disposed on the adhesion promoting layer is the analyte modulating layer in the form of a hydrophilic glucose limiting membrane comprising PDMS which modulates the diffusion of glucose therethrough.
  • a subsequent layer is a cover layer, typically composed of silicone, which is disposed on at least a portion of the analyte modulating layer, wherein the cover layer further includes an aperture that exposes at least a portion of the analyte modulating layer to the external glucose containing environment so that the glucose can access the analyte sensing layer on the working electrode.
  • This peroxide sensor apparatus functions via anodic polarization such that the hydrogen peroxide signal that is generated by glucose diffusing through the analyte modulating layer and then reacts with the glucose oxidase in the analyte sensing layer creates a detectable change in the current at the anodic working electrode in the conductive layer of the sensor that can be measured by an amperometer. This change in the current at the anodic working electrode can then be correlated with the concentration of glucose in the external environment. Consequently, a sensor of this design can act as a peroxide based glucose sensor.
  • the invention disclosed herein includes a number of embodiments including sensors having very thin enzyme coatings. Such embodiments of the invention allow artisans to generate a variety of permutations of the analyte sensor apparatus disclosed herein.
  • illustrative general embodiments of the sensor disclosed herein include a base layer, a cover layer and at least one layer having a sensor element such as an electrode disposed between the base and cover layers.
  • a sensor element such as an electrode disposed between the base and cover layers.
  • an exposed portion of one or more sensor elements e.g., a working electrode, a counter electrode, reference electrode, etc.
  • a very thin layer of material having an appropriate electrode chemistry.
  • FIG. 2 illustrates a cross-section of a typical sensor structure 100 of the present invention.
  • the sensor is formed from a plurality of layers of various conductive and non-conductive constituents disposed on each other according to a method of the invention to produce a sensor structure 100 .
  • the various layers (e.g. the analyte sensing layer) of the sensors can have one or more bioactive and/or inert materials incorporated therein.
  • incorporated as used herein is meant to describe any state or condition by which the material incorporated is held on the outer surface of or within a solid phase or supporting matrix of the layer.
  • the material “incorporated” may, for example, be immobilized, physically entrapped, attached covalently to functional groups of the matrix layer(s).
  • any process, reagents, additives, or molecular linker agents which promote the “incorporation” of said material may be employed if these additional steps or agents are not detrimental to, but are consistent with the objectives of the present invention.
  • a bioactive molecule e.g. an enzyme such as glucose oxidase
  • certain layers of the sensors disclosed herein include a proteinaceous substance such as albumin which serves as a crosslinkable matrix.
  • a proteinaceous substance is meant to encompass substances which are generally derived from proteins whether the actual substance is a native protein, an inactivated protein, a denatured protein, a hydrolyzed species, or a derivatized product thereof.
  • suitable proteinaceous materials include, but are not limited to enzymes such as glucose oxidase and lactate oxidase and the like, albumins (e.g. human serum albumin, bovine serum albumin etc.), caseins, gamma-globulins, collagens and collagen derived products (e.g., fish gelatin, fish glue, animal gelatin, and animal glue).
  • FIG. 2 An illustrative embodiment of the invention is shown in FIG. 2 .
  • This embodiment includes an electrically insulating base layer 102 to support the sensor 100 .
  • the electrically insulating layer base 102 can be made of a material such as a ceramic substrate, which may be self-supporting or further supported by another material as is known in the art.
  • the electrically insulating layer 102 comprises a polyimide substrate, for example a polyimide tape, dispensed from a reel. Providing the layer 102 in this form can facilitate clean, high density mass production. Further, in some production processes using such a polyimide tape, sensors 100 can be produced on both sides of the tape.
  • Typical embodiments of the invention include an analyte sensing layer disposed on the base layer 102 .
  • the analyte sensing layer comprises a conductive layer 104 which is disposed on insulating base layer 102 .
  • the conductive layer 104 comprises one or more electrodes.
  • the conductive layer 104 can be applied using many known techniques and materials as will be described hereafter, however, the electrical circuit of the sensor 100 is typically defined by etching the disposed conductive layer 104 into a desired pattern of conductive paths.
  • a typical electrical circuit for the sensor 100 comprises two or more adjacent conductive paths with regions at a proximal end to form contact pads and regions at a distal end to form sensor electrodes.
  • An electrically insulating protective cover layer 106 such as a polymer coating is typically disposed on portions of the conductive layer 104 .
  • Acceptable polymer coatings for use as the insulating protective layer 106 can include, but are not limited to, non-toxic biocompatible polymers such as polyimide, biocompatible solder masks, epoxy acrylate copolymers, or the like. Further, these coatings can be photo-imageable to facilitate photolithographic forming of apertures 108 through to the conductive layer 104 .
  • an analyte sensing layer is disposed upon a porous metallic and/or ceramic and/or polymeric matrix with this combination of elements functioning as an electrode in the sensor.
  • one or more exposed regions or apertures 108 can be made through the protective layer 106 to the conductive layer 104 to define the contact pads and electrodes of the sensor 100 .
  • the apertures 108 can be formed by a number of techniques, including laser ablation, chemical milling or etching or the like.
  • a secondary photoresist can also be applied to the cover layer 106 to define the regions of the protective layer to be removed to form the apertures 108 .
  • An operating sensor 100 typically includes a plurality of electrodes such as a working electrode and a counter electrode electrically isolated from each other, however typically situated in close proximity to one another. 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 can also undergo secondary processing through the apertures 108 , such as additional plating processing, to prepare the surfaces and/or strengthen the conductive regions.
  • An analyte sensing layer 110 is typically disposed on one or more of the exposed electrodes of the conductive layer 104 through the apertures 108 .
  • the analyte sensing layer 110 is a sensor chemistry layer and most typically an enzyme layer.
  • the analyte sensing layer 110 comprises the enzyme glucose oxidase or the enzyme lactate oxidase.
  • the analyte sensing layer 110 reacts with glucose to produce hydrogen peroxide which modulates a current to the electrode which can be monitored to measure an amount of glucose present.
  • the sensor chemistry layer 110 can be applied over portions of the conductive layer or over the entire region of the conductive layer.
  • the sensor chemistry layer 110 is disposed on portions of a working electrode and a counter electrode that comprise a conductive layer.
  • Some methods for generating the thin sensor chemistry layer 110 include spin coating processes, dip and dry processes, low shear spraying processes, ink-jet printing processes, silk 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 coating layers.
  • one such coating layer includes a membrane which can regulate the amount of analyte that can contact an enzyme of the analyte sensing layer.
  • a coating layer can comprise an analyte modulating membrane layer such as a glucose limiting membrane which regulates the amount of glucose that contacts the glucose oxidase enzyme layer on an electrode.
  • glucose limiting membranes can be made from a wide variety of materials known to be suitable for such purposes, e.g., silicone, polyurethane, polyurea cellulose acetate, Nafion, polyester sulfonic acid (Kodak AQ), hydrogels or any other membrane known to those skilled in the art.
  • materials known to be suitable for such purposes e.g., silicone, polyurethane, polyurea cellulose acetate, Nafion, polyester sulfonic acid (Kodak AQ), hydrogels or any other membrane known to those skilled in the art.
  • a coating layer is a glucose limiting membrane layer 112 which is disposed above the sensor chemistry layer 110 to regulate glucose contact with the sensor chemistry layer 110 .
  • an adhesion promoter layer 114 is disposed between the membrane layer 112 and the sensor chemistry layer 110 as shown in FIG. 2 in order to facilitate their contact and/or adhesion.
  • the adhesion promoter layer 114 can be made from any one of a wide variety of materials known in the art to facilitate the bonding between such layers.
  • the adhesion promoter layer 114 comprises a silane compound.
  • protein or like molecules in the sensor chemistry layer 110 can be sufficiently crosslinked or otherwise prepared to allow the membrane layer 112 to be disposed in direct contact with the sensor chemistry layer 110 in the absence of an adhesion promoter layer 114 .
  • embodiments of the present invention can include one or more functional coating layers.
  • the term “functional coating layer” denotes a layer that coats at least a portion of at least one surface of a sensor, more typically substantially all of a surface of the sensor, and that is capable of interacting with one or more analytes, such as chemical compounds, cells and fragments thereof, etc., in the environment in which the sensor is disposed.
  • functional coating layers include sensor chemistry layers (e.g., enzyme layers), analyte limiting layers, biocompatible layers; layers that increase the slipperiness of the sensor; layers that promote cellular attachment to the sensor; layers that reduce cellular attachment to the sensor; and the like.
  • analyte modulating layers operate to prevent or restrict the diffusion of one or more analytes, such as glucose, through the layers.
  • analytes such as glucose
  • Such layers can be formed to prevent or restrict the diffusion of one type of molecule through the layer (e.g. glucose), while at the same time allowing or even facilitating the diffusion of other types of molecules through the layer (e.g. O 2 ).
  • An illustrative functional coating layer is a hydrogel such as those disclosed in U.S. Pat. Nos. 5,786,439 and 5,391,250, the disclosures of each being incorporated herein by reference. The hydrogels described therein are particularly useful with a variety of implantable devices for which it is advantageous to provide a surrounding water layer.
  • the sensor embodiments disclosed herein can include layers having UV-absorbing polymers.
  • a sensor including at least one functional coating layer including an UV-absorbing polymer.
  • the UV-absorbing polymer is a polyurethane, a polyurea or a polyurethane/polyurea copolymer. More typically, the selected UV-absorbing polymer is formed from a reaction mixture including a diisocyanate, at least one diol, diamine or mixture thereof, and a polyfunctional UV-absorbing monomer.
  • UV-absorbing polymers are used with advantage in a variety of sensor fabrication methods, such as those described in U.S. Pat. No. 5,390,671, to Lord et al., entitled “Transcutaneous Sensor Insertion Set”; No. 5,165,407, to Wilson et al., entitled “Implantable Glucose Sensor”; and U.S. Pat. No. 4,890,620, to Gough, entitled “Two-Dimensional Diffusion Glucose Substrate Sensing Electrode”, which are incorporated herein in their entireties by reference.
  • any sensor production method which includes the step of forming an UV-absorbing polymer layer above or below a sensor element is considered to be within the scope of the present invention.
  • the inventive methods are not limited to thin-film fabrication methods, and can work with other sensor fabrication methods that utilize UV-laser cutting.
  • Embodiments can work with thick-film, planar or cylindrical sensors and the like, and other sensor shapes requiring laser cutting.
  • each sensor comprises a plurality of sensor elements, for example electrically conductive elements such as elongated thin film conductors, formed between an underlying insulative thin film base layer and an overlying insulative thin film cover layer.
  • electrically conductive elements such as elongated thin film conductors
  • a plurality of different sensor elements can be included in a single sensor.
  • both conductive and reactive sensor elements can be combined in one sensor, optionally with each sensor element being disposed on a different portion of the base layer.
  • One or more control elements can also be provided.
  • the sensor can have defined in its cover layer a plurality of openings or apertures.
  • One or more openings can also be defined in the cover layer directly over a portion of the base layer, in order to provide for interaction of the base layer with one or more analytes in the environment in which the sensor is disposed.
  • the base and cover layers can be comprised of a variety of materials, typically polymers. In more specific embodiments the base and cover layers are comprised of an insulative material such as a polyimide.
  • Openings are typically formed in the cover layer to expose distal end electrodes and proximal end contact pads.
  • the sensor can be placed transcutaneously so that the distal end electrodes are in contact with patient blood or extracellular fluid, and the contact pads are disposed externally for convenient connection to a monitoring device.
  • the sensors of the invention can have any desired configuration, for example planar or cylindrical.
  • the base layer 102 can be self-supportive, such as a rigid polymeric layer, or non-self supportive, such as a flexible film.
  • the latter embodiment is desirable in that it permits continuous manufacture of sensors using, for example, a roll of a polymeric film which is continuously unwound and upon which sensor elements and coating layers are continuously applied.
  • a general embodiment of the invention is a sensor designed for implantation within a body that comprises a base layer, an analyte sensing layer disposed upon the base layer which includes a plurality of sensor elements, an enzyme layer (typically less than 2 microns in thickness) disposed upon the analyte sensing layer which coats all of the plurality of sensing elements on the conductive layer, and one or more coating layers.
  • the enzyme layer comprises glucose oxidase; typically in a substantially fixed ratio with a carrier protein.
  • the glucose oxidase and the carrier protein are distributed in a substantially uniform manner throughout the disposed enzyme layer.
  • the carrier protein comprises albumin, typically in an amount of about 5% by weight.
  • a coating layer is an analyte contacting layer which is disposed on the sensor so as to regulate the amount of analyte that can contact the enzyme layer.
  • the sensor includes an adhesion promoter layer disposed between the enzyme layer and the analyte contacting layer; and, the enzyme layer is less than 1, 0.5, 0.25 or 0.1 microns in thickness.
  • One aspect of the present invention involves processes for making sensors having improved electrode chemistry coatings (e.g., enzyme coatings of less than 2 microns in thickness) with enhanced material properties.
  • Methods for producing the extremely thin enzyme coatings of the invention include spin coating processes, dip and dry processes, low shear spraying processes, ink-jet printing processes, silk screen processes and the like. Typically, such coatings are vapor crosslinked subsequent to their application.
  • sensors produced by these processes have material properties that exceed those of sensors having coatings produced by electrodeposition including enhanced longevity, linearity, regularity as well as improved signal to noise ratios.
  • certain sensor embodiments of the invention that utilize glucose oxidase coatings formed by such processes are designed to recycle hydrogen peroxide and improve the biocompatibility profiles of such sensors.
  • Illustrative embodiments of the invention include those designed to both consume hydrogen peroxide and recycle oxygen.
  • an illustrative embodiment of the invention is a method of making a less than about 2 micron coating of stabilized glucose oxidase on the surface of a matrix such as an electrode comprising combining glucose oxidase with albumin in a fixed ratio (one that is typically optimized for glucose oxidase stabilizing properties) and applying the glucose oxidase and albumin mixture to the surface of the matrix by a process selected from the group consisting of a spin coating process, a dip and dry process, a microdeposition process, a jet printer deposition process, a screen printing process or a doctor blading process.
  • a spin coating process Typically the stabilized glucose oxidase coating is applied to the surface of an electrode by a spin coating process.
  • the glucose oxidase/albumin is prepared in a physiological solution (e.g., phosphate buffered saline at neutral pH) with the albumin being present in an amount of about 5% albumin by weight.
  • a physiological solution e.g., phosphate buffered saline at neutral pH
  • albumin being present in an amount of about 5% albumin by weight.
  • the stabilized glucose oxidase layer that is formed on the conductive layer is less than 2, 1, 0.5, 0.25 or 0.1 microns in thickness.
  • a closely related embodiment of the invention is a stabilized glucose oxidase layer for coating the surface of an electrode wherein the glucose oxidase is mixed with a carrier protein in a fixed ratio within the layer, the glucose oxidase and the carrier protein are distributed in a substantially uniform manner throughout the layer. Typically the layer is less than 2 microns in thickness.
  • Embodiments of the invention include a design where an analyte sensing layer is disposed upon a porous metallic and/or ceramic and/or polymeric matrix with this combination of elements functioning as an electrode in the sensor.
  • a related embodiment of the invention is an electrochemical analyte sensor which includes a base layer, a conductive layer disposed upon the base layer that includes at least one working electrode and at least one counter electrode, an analyte sensing layer disposed upon the conductive layer, wherein the analyte sensing layer is less than 2 microns in thickness; and an analyte modulating layer that regulates the amount of analyte that contacts the enzyme layer, typically by limiting the amount of analyte that can diffuse through the layer and contact the analyte sensing layer.
  • the working electrode and/or the coated surface of the working electrode is larger than counter electrode and/or the coated surface of the counter electrode.
  • the enzyme layer comprises glucose oxidase stabilized by coating it on the working electrode and the counter electrode in combination with a carrier protein in a fixed ratio. In one embodiment, this glucose oxidase enzyme layer substantially covers the conductive layer. Embodiments where the glucose oxidase enzyme layer is disposed in a uniform coating over the whole conductive layer are typical because they may avoid problems associated with sensors having multiple different coatings on a single layer such as the selective delamination of different coatings having different material properties.
  • the sensor includes an adhesion promoting layer disposed between the enzyme layer and the analyte modulating layer.
  • a related embodiment of the invention is an electrochemical analyte sensor which includes a base layer, a conductive layer disposed upon the base layer that includes at least one working electrode, at least one reference electrode and at least one counter electrode, an enzyme layer disposed upon the conductive layer, and an analyte modulating cover layer that regulates the amount of analyte that contacts the enzyme layer.
  • the enzyme layer is less than 2 microns in thickness and is coated on at least a portion of the working electrode, the reference electrode and the counter electrode.
  • the enzyme layer substantially covers the working electrode, the reference electrode and the counter electrode.
  • the enzyme layer comprises glucose oxidase in combination with a carrier protein (e.g. albumin) in a fixed ratio.
  • the sensor includes an adhesion promoting layer disposed between the enzyme layer and the analyte modulating layer.
  • Yet another embodiment of the invention comprises a glucose sensor for implantation within a body which includes a base layer, a conductive layer disposed upon the base layer, an analyte sensing layer comprising glucose oxidase disposed upon the conductive layer, wherein the glucose oxidase is stabilized by combining it with albumin in a defined ratio and further wherein the glucose oxidase and the albumin are distributed in a substantially uniform manner throughout the disposed layer, and a glucose limiting layer that regulates the amount of glucose that diffuses through the glucose limiting layer and contacts the glucose oxidase layer.
  • the conductive layer includes a plurality of sensor elements including at least one working electrode and at least one counter electrode.
  • the analyte sensing layer comprising glucose oxidase is typically less than 2, 1, 0.5, 0.25 or 0.1 microns in thickness and the albumin in the layer is present in an amount of about 5% albumin by weight.
  • the sensor includes an adhesion promoting layer disposed between the analyte sensing layer comprising glucose oxidase and the glucose limiting layer.
  • planar refers to the well-known procedure of fabricating a substantially planar structure comprising layers of relatively thin materials, for example, using the well-known thick or thin-film techniques. See, for example, Liu et al., U.S. Pat. No. 4,571,292, and Papadakis et al., U.S. Pat. No. 4,536,274, both of which are incorporated herein by reference.
  • embodiments of the invention disclosed herein have a wider range of geometrical configurations (e.g. planar) than existing sensors in the art.
  • certain embodiments of the invention include one or more of the sensors disclosed herein coupled to another apparatus such as a medication infusion pump.
  • FIG. 2 provides a diagrammatic view of a typical analyte sensor configuration of the current invention.
  • Certain sensor configurations are of a relatively flat “ribbon” type configuration that can be made with the analyte sensor apparatus.
  • Such “ribbon” type configurations illustrate an advantage of the sensors disclosed herein that arises due to the spin coating of sensing enzymes such as glucose oxidase, a manufacturing step that produces extremely thin enzyme coatings that allow for the design and production of highly flexible sensor geometries.
  • Such thin enzyme coated sensors provide further advantages such as allowing for a smaller sensor area while maintaining sensor sensitivity, a highly desirable feature for implantable devices (e.g. smaller devices are easier to implant).
  • sensor embodiments of the invention that utilize very thin analyte sensing layers that can be formed by processes such as spin coating can have a wider range of geometrical configurations (e.g. planar) than those sensors that utilize enzyme layers formed via processes such as electrodeposition.
  • Certain sensor configurations include multiple conductive elements such as multiple working, counter and reference electrodes. Advantages of such configurations include increased surface area which provides for greater sensor sensitivity. For example, one sensor configuration introduces a third working sensor. One obvious advantage of such a configuration is signal averaging of three sensors which increases sensor accuracy. Other advantages include the ability to measure multiple analytes. In particular, analyte sensor configurations that include electrodes in this arrangement (e.g. multiple working, counter and reference electrodes) can be incorporated into multiple analyte sensors.
  • analytes such as oxygen, hydrogen peroxide, glucose, lactate, potassium, calcium, and any other physiologically relevant substance/analyte provides a number of advantages, for example the ability of such sensors to provide a linear response as well as ease in calibration and/or recalibration.
  • An exemplary multiple sensor device comprises a single device having a first sensor which is polarized cathodically and designed to measure the changes in oxygen concentration that occur at the working electrode (a cathode) as a result of glucose interacting with glucose oxidase; and a second sensor which is polarized anodically and designed to measure changes in hydrogen peroxide concentration that occurs at the working electrode (an anode) as a result of glucose coming form the external environment and interacting with glucose oxidase.
  • the first oxygen sensor will typically experience a decrease in current at the working electrode as oxygen contacts the sensor while the second hydrogen peroxide sensor will typically experience an increase in current at the working electrode as the hydrogen peroxide generated as shown in FIG. 1 contacts the sensor.
  • an observation of the change in current that occurs at the working electrodes as compared to the reference electrodes in the respective sensor systems correlates to the change in concentration of the oxygen and hydrogen peroxide molecules which can then be correlated to the concentration of the glucose in the external environment (e.g. the body of the mammal).
  • the analyte sensors of the invention can be coupled with other medical devices such as medication infusion pumps.
  • replaceable analyte sensors of the invention can be coupled with other medical devices such as medication infusion pumps, for example by the use of a port couple to the medical device (e.g. a subcutaneous port with a locking electrical connection).
  • Typical sensors for monitoring glucose concentration of diabetics are further described in Shictiri, et al.,: “In Vivo Characteristics of Needle-Type Glucose Sensor-Measurements of Subcutaneous Glucose Concentrations in Human Volunteers,” Horm. Metab. Res., Suppl. Ser. 20:17-20 (1988); Bruckel, et al.,: “In Vivo Measurement of Subcutaneous Glucose Concentrations with an Enzymatic Glucose Sensor and a Wick Method,” Klin. Weinschr.
  • a typical embodiment of the invention disclosed herein is a method of making a sensor apparatus for implantation within a mammal comprising the steps of: providing a base layer; forming a conductive layer on the base layer, wherein the conductive layer includes an electrode (and typically a working electrode, a reference electrode and a counter electrode); forming an analyte sensing layer on the conductive layer, wherein the analyte sensing layer includes a composition that can alter the electrical current at the electrode in the conductive layer in the presence of an analyte; optionally forming a protein layer on the analyte sensing layer; forming an adhesion promoting layer on the analyte sensing layer or the optional protein layer; forming an analyte modulating layer disposed on the adhesion promoting layer, wherein the analyte modulating layer includes a composition that modulates the diffusion of the analyte therethrough; and forming a cover layer disposed on at least a portion of the analyte
  • the various layers of the sensor can be manufactured to exhibit a variety of different characteristics which can be manipulated according to the specific design of the sensor.
  • the adhesion promoting layer includes a compound selected for its ability to stabilize the overall sensor structure, typically a silane composition.
  • the analyte sensing layer is formed by a spin coating process and is of a thickness selected from the group consisting of less than 1, 0.5, 0.25 and 0.1 microns in height.
  • a method of making the sensor includes the step of forming a protein layer on the analyte sensing layer, wherein a protein within the protein layer is an albumin selected from the group consisting of bovine serum albumin and human serum albumin.
  • a method of making the sensor includes the step of forming an analyte sensing layer that comprises an enzyme composition selected from the group consisting of glucose oxidase, glucose dehydrogenase, lactate oxidase, hexokinase and lactate dehydrogenase.
  • the analyte sensing layer typically comprises a carrier protein composition in a substantially fixed ratio with the enzyme and the enzyme and the carrier protein are distributed in a substantially uniform manner throughout the analyte sensing layer.
  • One embodiment of the invention comprises porous metallic matrices.
  • metallic substrate embodiments of the invention can be manufactured with the desired porosity, pore-size distribution, and tortuosity through a printing process.
  • the metallic substrate may either be printed as a film or within the confines of a mold, either directly in place onto the sensor assembly or onto a temporary substrate.
  • the ink can consist of fine metallic particles suspended in a porogenic carrier.
  • the metallic particles may consist of a single pure metal or alloy. Different types of metallic particles may also be printed either at the same time to form a mixture, or at different times to form layers.
  • the porogenic carrier can consist of a solvent with or without various polymers, glasses, ceramics, and/or frit materials.
  • the mold may consist of various ceramics, polymers, or metals.
  • the printed metallic matrix can be dried at an appropriate temperature.
  • the resulting porous bed of metallic powder can then be fired approximately in the range of 350° C.-2,000° C. to bond the metallic and, if any, ceramic particles together.
  • This can form a highly porous and tortuous metallic substrate onto which an enzyme such as GOx can be immobilized.
  • the morphology of the metallic substrate can be adjusted by manipulating the size of the metallic particles as well as the composition of the porogenic carrier.
  • various glass, ceramic, and/or metallic particles included in the ink can be etched from the printed material to create pores using materials such as, but not limited to, hydrofluoric acid and sodium hydroxide.
  • materials such as, but not limited to, hydrofluoric acid and sodium hydroxide.
  • platinum black Prior to coating the metallic substrate with glucose oxidase, platinum black may or may not be plated using standard techniques.
  • the matrix may comprise a metallic foam.
  • Porous metals are those that contain a multitude of pores, i.e. closed, curved gas voids with a smooth surface.
  • Metal(lic) foams are special cases of porous metals.
  • a solid foam originates from a liquid foam in which gas bubbles are finely dispersed in a liquid.
  • space is filled by pieces of metal that form a continuous network and co-exist with a network of empty space which is also interconnected.
  • the porous metallic substrates can be manufactured by drilling small holes into a metal sheet, film, foil, rod, or block using a laser beam or some other type of drilling technology.
  • a woven wire mesh can be used as a porous metallic substrate.
  • the fabrication of 3-D micromesh Ni Structures using electroplating has been described in the art such as fabrication methods of a 3-D micromesh Ni electrode.
  • inverse-micromesh photoresist structures fabricated by multiple inclined backside exposure, can be used as a mold for Ni electroplating, with Ni meshes of about 3 ⁇ m in diameter obtained by this method.
  • the enzyme composition can be applied to the porous matrices by any one of a variety of methods known in the art.
  • an enzyme such as glucose oxidase can be dissolved in a solvent and dip, spray, or spin coated onto the porous metallic substrate.
  • the coating solvent may consist of aqueous buffer and/or various organic solvents and/or surfactants including, but not limited to, various alcohols, dimethyl sulfoxide, and polyoxyethylene (20) sorbitan monolaurate (“TweenTM 20”).
  • Ingress of the protein into porous substrates may be promoted by decreasing the viscosity of the enzyme solution through the manipulation of its composition and/or by applying vacuum and/or centrifugation and/or ultrasonic vibration to the coated substrate.
  • Other bio and/or synthetic polymers may also be coated along with the enzyme as filler material such as, but not limited to: bovine serum albumin, human serum albumin, polyethylene glycol, and O′,O′-Bis(2-aminopropyl)polyethylene glycol (“Jeffamine®”).
  • the coated enzyme and filler materials will be immobilized onto the metallic substrate using an appropriate homobifunctional (i.e. glutaraldehyde or disuccinimidyl suberate), heterobifunctional (i.e.
  • PHOTOLINKTM a process provided by SurModics Inc. under the trademark PHOTOLINKTM can be used to immobilize an enzyme such as glucose oxidase onto the porous metallic substrate.
  • PHOTOLINKTM methods are set forth in U.S. Pat. Nos. 3,959,078, 4,722,906, 5,229,172; 5,308,641; 5,350,800 and 5,415,938.
  • embodiments of the invention include an essentially rigid, non-swelling porous enzyme-polymer matrix.
  • molded continuous rods of macroporous polymers have been developed for use as chromatographic separation media (see, e.g. U.S. Pat. No. 5,453,185 and PCT Publication No. WO 93/07945, the contents of which are incorporated herein by reference). Examples include, but are not limited to poly(glycidyl methacrylate-co-ethylene dimethacrylate) and poly(styrene-co-divinylbenzene). As disclosed in U.S. Pat. No.
  • a typical polymerization mixture at a minimum contains at least one polyvinyl monomer, a free radical generating initiator, and a porogen.
  • the mixture may also contain one or more monovinyl monomers and/or soluble polymers or insoluble macroporous polymer particles.
  • Suitable polyvinyl monomers include divinylbenzene, divinylnaphthalene, divinylpyridine, alkylene dimethacrylates, hydroxyalkylene dimethacrylates, hydroxyalkylene diacrylates, oligoethylene glycol dimethacrylates, oligoethylene glycol diacrylates, vinyl esters of polycarboxylic acids, divinyl ether, pentaerythritol di-, tri-, or tetramethacrylate or acrylate, trimethylopropane trimethacrylate or acrylate, alkylene bis acrylamides or methacrylamides, and mixtures of any such suitable polyvinyl monomers.
  • the alkylene groups generally contain about 1-6 carbon atoms.
  • Monovinyl monomers which may be used include styrene, ring substituted styrenes wherein the substituents include chloromethyl, alkyl with up to 18 carbon atoms, hydroxyl, t-butyloxycarbonyl, halogen, nitro, amino group, protected hydroxyls or amino groups, vinylnaphthalene, acrylates, methacrylates, vinylacetate, vinylpyrolidone, and mixtures thereof.
  • the polyvinyl monomer or polyvinyl monomer plus the monovinyl monomer are generally present in the polymerization mixture in an amount of from about 10 to 60 vol. %, and more typically in an amount of from about 20 to 40 vol. %.
  • the porogen that is used may be selected from a variety of different types of materials.
  • suitable liquid porogens include aliphatic hydrocarbons, aromatic hydrocarbons, esters, alcohols, ketones, ethers, solutions of soluble polymers, and mixtures thereof.
  • the porogen is generally present in the polymerization mixture in an amount of from about 40 to 90 vol %, more typically from about 60 to 80 vol %.
  • Soluble polymers and insoluble polymer particles may be employed in combination with the monomers. These polymers are added to the polymerization mixture prior to polymerization.
  • the soluble polymers are dissolved out of the plug after its formation by passing a solvent through the plug.
  • the soluble polymers serve as a polymeric porogen to increase the porosity of the final plug.
  • Suitable soluble polymers used herein include non-crosslinked polymers or copolymers of such monomers as styrene or ring substituted styrene, acrylates, methacrylates, dienes, vinylchloride, and vinylacetate.
  • the insoluble polymer particles are used to reduce the volume shrinkage during the polymerization. The lesser the volume of the monomers in the polymerization mixture the smaller the contraction of volume upon polymerization.
  • Suitable insoluble polymer particles used herein include macroporous polymer particles which are cross-linked copolymers of the same monomers. It is, however, common due to compatibility to employ insoluble polymer particles which are formed from the same monomers used to form the polymerization mixture with which they are to be combined.
  • the polymer particles initially have a diameter of from about 1 to 1,000 micrometers. It is not necessary that the mixture of polymer particles have the same particle size. In fact, it is more economical, and, therefore common to use irregularly sized polymer particles. While not necessary, the polymer particles may be soaked with a liquid immiscible with the polymerization mixture which can contain an inhibitor which inhibits free radical polymerization. This is done in order to prevent polymerization in the inside of the macroporous particles which would cause filling of the pores and would effectively remove them from the separation process. The rod would then contain nonporous pools unable to contribute to the separation process. Suitable inhibitors include cupric chloride and sodium nitrite.
  • the inhibitor is generally present in an amount of from about 0,001 to 1 wt %, and more typically in an amount of from about 0.1 to 1 wt %, based on the total weight of particles.
  • the polymer particles are typically degassed prior to use in the polymerization mixture. This may be accomplished by any of the conventional means known in the art. It, however, is typical to soak the particles in water, optionally containing a polymerization inhibitor, and remove the air from the pores by keeping the water-polymer particle mixture under the vacuum of a water pump for a suitable period of time such as about 5 to 20 minutes. Excess water may then be removed by filtering.
  • the soluble polymers are generally present in an amount of from about 5 to 40% by volume of the polymerization mixture and the insoluble polymer particles in an amount of from about 5 to 50% by volume.
  • Conventional free-radical generating polymerization initiators may be employed to initiate polymerization.
  • suitable initiators include peroxides such as OO-t-amyl-O-(2-ethylhexyl)monoperoxycarbonate, dipropylperoxydicarbonate, and benzoyl peroxide, as well as azo compounds such as azobisisobutyronitrile, 2,2′-azobis(2-amidinopropane)dihydrochloride, and 2,2′-azobis(isobutyramide)dihydrate. It has been found that the choice of initiator may be used as a means to control the pore distribution in a plug.
  • the initiator is generally present in the polymerization mixture in an amount of from about 0.2 to 5% by weight of the monomers.
  • Polymers useful for making the essentially rigid, non-swelling porous enzyme-polymer matrices are essentially incompressible and do not change their overall size in response to changes in their solvating environment. Adjustments to the polymerization conditions can be used to control the morphology of the pores. Hence, highly porous (50-70%) polymers can be created that possess significant volume fractions of pores in the ranges of 1-100 nm and 100-3,000 nm (i.e. 20% and 80%, respectively). Polymers with this type of pore structure possess a very high specific surface area (i.e. 185 m 2 /g), and are expected to allow for high enzyme immobilization densities (1-100 mg/mL).
  • a nucleophilic compound can be used to functionalize a macroporous, rigid polymer that possesses reactive epoxide groups.
  • a cross-linking agent can then be used to immobilize the bio-sensing enzyme to the polymer via the functional groups of the enzyme and polymer substrate.
  • Other nucleophilic compounds that can be used to functionalize epoxide-activated polymers include, but are not limited to ammonia, ethylenediamine, ethanolamine, carbohydrates, cysteine, and other amino acids. For a given enzyme and functionalized polymer combination, an appropriate homobifunctional (i.e.
  • disuccinimidyl suberate i.e. succinimidyl-4-[N-maleimidomethyl]cyclohexane-1-carboxylate
  • trifunctional i.e. 4-azido-2-nitrophenylbiocytin-4-nitrophenyl ester
  • zero-length i.e. 1-ethyl-3-[3-dimethylaminopropyl]carbodiimide hydrochloride
  • the biosensing enzyme will be directly immobilized onto an epoxide-activated polymer via nucleophilic attack by sulfhydryl, amine, hydroxyl, and/or carboxyl groups that are either native to the enzyme, or have been added to the wild-type peptide sequence via genetic engineering or directed evolution.
  • the nucleophilic functional groups of the enzyme may be reversibly or irreversibly blocked or protected during the immobilization, using compounds that would be familiar to anyone well versed in protein conjugation (i.e. 5,5′-dithio-bis-[2-nitrobenzoic acid] or N-ethylmaleimide).
  • rigid, non-swelling porous enzyme-polymer matrices monomers possessing functional groups other than (or in addition to) epoxide groups will be incorporated into the rigid, macroporous polymer during the polymerization reaction (i.e. aminostyrene).
  • the bio-sensing enzyme could then be immobilized onto the polymer substrate using an appropriate homobifunctional, heterobifunctional, trifunctional, and/or zero-length cross-linking agent.
  • PhotoLink® (SurModics, Eden Prairie, Minn.) chemistry can be used to immobilize the bio-sensing enzyme to the molded, porous, rigid polymer.
  • the polymer substrate need not possess any functional groups because the PhotoLink® chemistry reacts with carbon-hydrogen groups found in virtually every organic polymer.
  • the disclosure provided herein includes sensors and sensor designs that can be generated using combinations of various well known techniques.
  • the disclosure further provides methods for applying very thin enzyme coatings to these types of sensors as well as sensors produced by such processes.
  • some embodiments of the invention include methods for making such sensors on a substrate according to art accepted processes.
  • the substrate comprises a rigid and flat structure suitable for use in photolithographic mask and etch processes.
  • the substrate typically defines an upper surface having a high degree of uniform flatness.
  • a polished glass plate may be used to define the smooth upper surface.
  • Alternative substrate materials include, for example, stainless steel, aluminum, and plastic materials such as delrin, etc.
  • the substrate is non-rigid and can be another layer of film or insulation that is used as a substrate, for example plastics such as polyimides and the like.
  • An initial step in the methods of the invention typically includes the formation of a base layer of the sensor.
  • the base layer can be disposed on the substrate by any desired means, for example by controlled spin coating.
  • an adhesive may be used if there is not sufficient adhesion between the substrate layer and the base layer.
  • a base layer of insulative material is formed on the substrate, typically by applying the base layer material onto the substrate in liquid form and thereafter spinning the substrate to yield the base layer of thin, substantially uniform thickness. These steps are repeated to build up the base layer of sufficient thickness, followed by a sequence of photolithographic and/or chemical mask and etch steps to form the conductors discussed below.
  • the base layer comprises a thin film sheet of insulative material, such as ceramic or polyimide substrate.
  • the base layer can comprise an alumina substrate, a polyimide substrate, a glass sheet, controlled pore glass, or a planarized plastic liquid crystal polymer.
  • the base layer may be derived from any material containing one or more of a variety of elements including, but not limited to, carbon, nitrogen, oxygen, silicon, sapphire, diamond, aluminum, copper, gallium, arsenic, lanthanum, neodymium, strontium, titanium, yttrium, or combinations thereof.
  • the substrate may be coated onto a solid support by a variety of methods well-known in the art including chemical vapor deposition, physical vapor deposition, or spin-coating with materials such as spin glasses, chalcogenides, graphite, silicon dioxide, organic synthetic polymers, and the like.
  • the methods of the invention further include the generation of a conductive layer having one or more sensing elements.
  • these sensing elements are electrodes that are formed by one of the variety of methods known in the art such as photoresist, etching and rinsing to define the geometry of the active electrodes.
  • the electrodes can then be made electrochemically active, for example by electrodeposition of Pt black for the working and counter electrode, and silver followed by silver chloride on the reference electrode.
  • a sensor layer such as a sensor chemistry enzyme layer can then be disposed on the sensing layer by electrochemical deposition or a method other than electrochemical deposition such a spin coating, followed by vapor crosslinking, for example with a dialdehyde (glutaraldehyde) or a carbodi-imide.
  • Electrodes of the invention can be formed from a wide variety of materials known in the art.
  • the electrode may be made of a noble late transition metals.
  • Metals such as gold, platinum, silver, rhodium, iridium, ruthenium, palladium, or osmium can be suitable in various embodiments of the invention.
  • Other compositions such as carbon or mercury can also be useful in certain sensor embodiments.
  • silver, gold, or platinum is typically used as a reference electrode metal.
  • a silver electrode which is subsequently chloridized is typically used as the reference electrode.
  • metals can be deposited by any means known in the art, including the plasma deposition method cited, supra, or by an electroless method which may involve the deposition of a metal onto a previously metallized region when the substrate is dipped into a solution containing a metal salt and a reducing agent.
  • the electroless method proceeds as the reducing agent donates electrons to the conductive (metallized) surface with the concomitant reduction of the metal salt at the conductive surface. The result is a layer of adsorbed metal.
  • the base layer is initially coated with a thin film conductive layer by electrode deposition, surface sputtering, or other suitable process step.
  • this conductive layer may be provided as a plurality of thin film conductive layers, such as an initial chrome-based layer suitable for chemical adhesion to a polyimide base layer followed by subsequent formation of thin film gold-based and chrome-based layers in sequence.
  • other electrode layer conformations or materials can be used.
  • the conductive layer is then covered, in accordance with conventional photolithographic techniques, with a selected photoresist coating, and a contact mask can be applied over the photoresist coating for suitable photoimaging.
  • the contact mask typically includes one or more conductor trace patterns for appropriate exposure of the photoresist coating, followed by an etch step resulting in a plurality of conductive sensor traces remaining on the base layer.
  • each sensor trace can include three parallel sensor elements corresponding with three separate electrodes such as a working electrode, a counter electrode and a reference electrode.
  • an insulative cover layer typically of a material such as a silicon polymer and/or a polyimide.
  • the insulative cover layer can be applied in any desired manner.
  • the insulative cover layer is applied in a liquid layer over the sensor traces, after which the substrate is spun to distribute the liquid material as a thin film overlying the sensor traces and extending beyond the marginal edges of the sensor traces in sealed contact with the base layer.
  • This liquid material can then be subjected to one or more suitable radiation and/or chemical and/or heat curing steps as are known in the art.
  • the liquid material can be applied using spray techniques or any other desired means of application.
  • Various insulative layer materials may be used such as photoimagable epoxyacrylate, with an illustrative material comprising a photoimagable polyimide available from OCG, Inc. of West Paterson, N.J., under the product number 7020.
  • appropriate electrode chemistries defining the distal end electrodes can be applied to the sensor tips, optionally subsequent to exposure of the sensor tips through the openings.
  • an enzyme typically glucose oxidase
  • an enzyme is provided within one of the openings, thus coating one of the sensor tips to define a working electrode.
  • One or both of the other electrodes can be provided with the same coating as the working electrode.
  • the other two electrodes can be provided with other suitable chemistries, such as other enzymes, left uncoated, or provided with chemistries to define a reference electrode and a counter electrode for the electrochemical sensor.
  • a significant aspect of the present invention involves processes for making sensors having extremely thin coatings for electrode chemistries (e.g., enzyme coatings of less than 2 microns in thickness) with enhanced material properties.
  • Methods for producing the extremely thin enzyme coatings of the invention include spin coating processes, dip and dry processes, low shear spraying processes, ink-jet printing processes, silk screen processes and the like.
  • artisans can readily determine the thickness of an enzyme coat applied by process of the art, they can readily identify those methods capable of generating the extremely thin coatings of the invention.
  • coatings are vapor crosslinked subsequent to their application.
  • sensors produced by these processes have material properties that exceed those of sensors having coatings produced by electrodeposition including enhanced longevity, linearity, regularity as well as improved signal to noise ratios.
  • embodiments of the invention that utilize glucose oxidase coatings formed by such processes are designed to recycle hydrogen peroxide and improve the biocompatibility profiles of such sensors.
  • the surprising properties of sensors produced by such processes have enhanced characteristics as compared to those generated by electrodeposition because electrodeposition produces 3-5 micron thick enzyme layers in which only a fraction of the reactive enzyme is able to access the analyte to be sensed.
  • the thick coatings produced by electrodeposition may hinder the ability of hydrogen peroxide generated at the reactive interface to reach the sensor surface and thereby generate a signal.
  • hydrogen peroxide that is unable to reach a sensor surface due to such thick coatings typically diffuses away from the sensor into the environment in which the sensor is placed, thereby decreasing the biocompatibility of such sensors.
  • electrodeposition processes can result in a surface coating in which an optimally determined ratio of enzyme to carrier protein is detrimentally altered and further wherein the glucose oxidase and the carrier protein are not distributed in a substantially uniform manner throughout the disposed enzyme layer.
  • the thin coating processes utilized to produce the sensors disclosed herein avoid these problems associated with electrodeposition.
  • Sensors generated by processes such as spin coating processes also avoid other problems associated with electrodeposition, such as those pertaining to the material stresses placed on the sensor during the electrodeposition process.
  • the process of electrodeposition is observed to produce mechanical stresses on the sensor, for example mechanical stresses that result from tensile and/or compression forces.
  • mechanical stresses may result in sensors having coatings with some tendency to crack or delaminate. This is not observed in coatings disposed on sensor via spin coating or other low-stress processes. Consequently, yet another embodiment of the invention is a method of avoiding the electrodeposition influenced cracking and or delamination of a coating on a sensor comprising applying the coating via a spin coating process.
  • one or more additional functional coating or cover layers can then be applied by any one of a wide variety of methods known in the art, such as spraying, dipping, etc.
  • Some embodiments of the present invention include an analyte modulating layer deposited over the enzyme-containing layer.
  • an analyte modulating layer deposited over the enzyme-containing layer.
  • the thickness of the analyte modulating membrane layer can influence the amount of analyte that reaches the active enzyme.
  • nonsiloxane-siloxane copolymers include, but are not limited to, dimethylsiloxane-alkene oxide, tetramethyldisiloxane-divinylbenzene, tetramethyldisiloxane-ethylene, dimethylsiloxane-silphenylene, dimethylsiloxane-silphenylene oxide, dimethylsiloxane-a-methylstyrene, dimethylsiloxane-bisphenol A carbonate copolymers, or suitable combinations thereof.
  • the percent by weight of the nonsiloxane component of the copolymer can be preselected to any useful value but typically this proportion lies in the range of about 40-80 wt %.
  • the dimethylsiloxane-bisphenol A carbonate copolymer which comprises 50-55 wt % of the nonsiloxane component is typical. These materials may be purchased from Petrarch Systems, Bristol, Pa. (USA) and are described in this company's products catalog. Other materials which may serve as analyte limiting membrane layers include, but are not limited to, polyurethanes, cellulose acetate, cellulose nitrate, silicone rubber, or combinations of these materials including the siloxane nonsiloxane copolymer, where compatible.
  • the senor is made by methods which apply an analyte modulating layer that comprises a hydrophilic membrane coating which can regulate the amount of analyte that can contact the enzyme of the sensor layer.
  • the cover layer that is added to the glucose sensors of the invention can comprise a glucose limiting membrane, which regulates the amount of glucose that contacts glucose oxidase enzyme layer on an electrode.
  • glucose limiting membranes can be made from a wide variety of materials known to be suitable for such purposes, e.g., silicones such as polydimethyl siloxane and the like, polyurethanes, cellulose acetates, Nafion, polyester sulfonic acids (e.g.
  • the membrane layer that is disposed on the glucose oxidase enzyme layer functions to inhibit the release of hydrogen peroxide into the environment in which the sensor is placed and to facilitate the contact between the hydrogen peroxide molecules and the electrode sensing elements.
  • an adhesion promoter layer is disposed between a cover layer (e.g. an analyte modulating membrane layer) and a sensor chemistry layer in order to facilitate their contact and is selected for its ability to increase the stability of the sensor apparatus.
  • a cover layer e.g. an analyte modulating membrane layer
  • a sensor chemistry layer e.g. an analyte modulating membrane layer
  • compositions of the adhesion promoter layer are selected to provide a number of desirable characteristics in addition to an ability to provide sensor stability. For example, some compositions for use in the adhesion promoter layer are selected to play a role in interference rejection as well as to control mass transfer of the desired analyte.
  • the adhesion promoter layer can be made from any one of a wide variety of materials known in the art to facilitate the bonding between such layers and can be applied by any one of a wide variety of methods known in the art.
  • the adhesion promoter layer comprises a silane compound such as ⁇ -aminopropyltrimethoxysilane.
  • the adhesion promoting layer and/or the analyte modulating layer comprises an agent selected for its ability to crosslink a siloxane moiety present in a proximal.
  • the adhesion promoting layer and/or the analyte modulating layer comprises an agent selected for its ability to crosslink an amine or carboxyl moiety of a protein present in a proximal layer.
  • the AP layer further comprises Polydimethyl Siloxane (PDMS), a polymer typically present in analyte modulating layers such as a glucose limiting membrane.
  • PDMS Polydimethyl Siloxane
  • the formulation comprises 0.5-20% PDMS, typically 5-15% PDMS, and most typically 10% PDMS.
  • the addition of PDMS to the AP layer can be advantageous in contexts where it diminishes the possibility of holes or gaps occurring in the AP layer as the sensor is manufactured.
  • a coupling reagent commonly used for promoting adhesion between sensor layers is ⁇ -aminopropyltrimethoxysilane.
  • the silane compound is usually mixed with a suitable solvent to form a liquid mixture.
  • the liquid mixture can then be applied or established on the wafer or planar sensing device by any number of ways including, but not limited to, spin-coating, dip-coating, spray-coating, and microdispensing.
  • the microdispensing process can be carried out as an automated process in which microspots of material are dispensed at multiple preselected areas of the device.
  • Solvents suitable for use in forming the silane mixtures include aqueous as well as water-miscible organic solvents, and mixtures thereof. Alcoholic water-miscible organic solvents and aqueous mixtures thereof are particularly useful. These solvent mixtures may further comprise nonionic surfactants, such as polyethylene glycols (PEG) having a for example a molecular weight in the range of about 200 to about 6,000.
  • PEG polyethylene glycols
  • These hydrolyzed silane reagents are also able to condense with polar groups, such as hydroxyls, which may be present on the substrate surface.
  • polar groups such as hydroxyls, which may be present on the substrate surface.
  • atmospheric moisture may be sufficient to hydrolyze the alkoxy groups present initially on the silane reagent.
  • the R′ group usually contains a terminal amine group useful for the covalent attachment of an enzyme to the substrate surface (a compound, such as glutaraldehyde, for example, may be used as a linking agent as described by Murakami, T. et al., Analytical Letters 1986, 19, 1973-86).
  • a compound such as glutaraldehyde, for example, may be used as a linking agent as described by Murakami, T. et al., Analytical Letters 1986, 19, 1973-86).
  • the adhesion promoter layer can be subjected to one or more suitable radiation and/or chemical and/or heat curing steps as are known in the art.
  • the enzyme layer can be sufficiently crosslinked or otherwise prepared to allow the membrane cover layer to be disposed in direct contact with the sensor chemistry layer in the absence of an adhesion promoter layer.
  • An illustrative embodiment of the invention is a method of making a sensor by providing a base layer, forming a sensor layer on the base layer, spin coating an enzyme layer on the sensor layer and then forming an analyte contacting layer (e.g. an analyte modulating layer such as a glucose limiting membrane) on the sensor, wherein the analyte contacting layer regulates the amount of analyte that can contact the enzyme layer.
  • the enzyme layer is vapor crosslinked on the sensor layer.
  • the sensor layer is formed to include at least one working electrode and at least one counter electrode.
  • the enzyme layer is formed on at least a portion of the working electrode and at least a portion of the counter electrode.
  • the enzyme layer that is formed on the sensor layer is less than 2, 1, 0.5, 0.25 or 0.1 microns in thickness.
  • the enzyme layer comprises one or more enzymes such as glucose oxidase, glucose dehydrogenase, lactate oxidase, hexokinase or lactate dehydrogenase and/or like enzymes.
  • the enzyme layer comprises glucose oxidase that is stabilized by coating it on the sensor layer in combination with a carrier protein in a fixed ratio.
  • the carrier protein is albumin.
  • such methods include the step of forming an adhesion promoter layer disposed between the glucose oxidase layer and the analyte contacting layer.
  • the adhesion promoter layer is subjected to a curing process prior to the formation of the analyte contacting layer.
  • a related embodiment of the invention is a method of making a glucose sensor by providing a base layer, forming a sensor layer on the base layer that includes at least one working electrode and at least one counter electrode, forming a glucose oxidase layer on the sensor layer by a spin coating process (a layer which is typically stabilized by combining the glucose oxidase with albumin in a fixed ratio), wherein the glucose oxidase layer coats at least a portion of the working electrode and at least a portion of the counter electrode, and then forming a glucose limiting layer on the glucose sensor so as to regulate the amount of glucose that can contact the glucose oxidase layer.
  • a spin coating process a layer which is typically stabilized by combining the glucose oxidase with albumin in a fixed ratio
  • the glucose oxidase layer that is formed on the sensor layer is typically less than 2, 1, 0.5, 0.25 or 0.1 microns in thickness.
  • the glucose oxidase coating is vapor crosslinked on the sensor layer.
  • the glucose oxidase coating covers the entire sensor layer.
  • an adhesion promoter layer is disposed between the glucose oxidase layer and the analyte contacting layer.
  • the analyte sensor further comprises one or more cover layers which are typically electrically insulating protective layers (see, e.g. element 106 in FIG. 2 ).
  • cover layers are disposed on at least a portion of the analyte modulating layer.
  • the finished sensors produced by such processes are typically quickly and easily removed from a supporting substrate (if one is used), for example, by cutting along a line surrounding each sensor on the substrate.
  • the cutting step can use methods typically used in this art such as those that include a UV laser cutting device that is used to cut through the base and cover layers and the functional coating layers along a line surrounding or circumscribing each sensor, typically in at least slight outward spaced relation from the conductive elements so that the sufficient interconnected base and cover layer material remains to seal the side edges of the finished sensor.
  • dicing techniques typically used to cut ceramic substrates can be used with the appropriate sensor embodiments.
  • the sensors can be lifted quickly and easily from the supporting substrate, without significant further processing steps or potential damage due to stresses incurred by physically pulling or peeling attached sensors from the supporting substrate.
  • the supporting substrate can thereafter be cleaned and reused, or otherwise discarded.
  • the functional coating layer(s) can be applied either before or after other sensor components are removed from the supporting substrate (e.g., by cutting).
  • Another illustrative embodiment of the invention is a method of making a metallic electrode useful in biosensors and electrodes made by such methods.
  • These methods for making a metallic electrode include electrodepositing a plurality of metal layers that comprise the electrode using cycles of differing electroplating conditions.
  • the method comprises a first cycle of electroplating where a metal is electrodeposited onto a substrate under a first set of conditions selected to produce a first metal layer having a first surface area and a first adhesion strength between the substrate and the first metal layer.
  • the method then involves a second cycle of electroplating where a metal composition is then electrodeposited onto the first metal layer under a second set of conditions selected to produce a second metal layer having a second surface area and a second adhesion strength between the first metal layer and the second metal layer.
  • the first and second set of conditions are selected to produce a second metal layer having a second surface area that is greater (e.g. at least 5%, 10%, 15%, 20%, 25%, or 30% greater) than the first surface area of the first metal layer produced by the first set of conditions and a second metal layer having an adhesion with the first metal layer that is greater (e.g.
  • the method further comprises additional cycles of electroplating.
  • the method comprises a third cycle of electroplating where a metal composition is electrodeposited onto the second layer under a third set of conditions selected to produce a third metal layer having a third surface area.
  • the second and third set of conditions are selected to produce a third metal layer having a greater density (e.g. at least 5%, 10%, 15%, 20%, 25%, or 30% greater) than the density of the second metal layer.
  • the third set of conditions produces a third metal layer having a third surface area that is less than the second surface area of the second metal layer produced by the second set of conditions.
  • the substrate does not comprise platinum and comprises a polymer or another metal such as gold.
  • the substrate comprises a geometric feature selected to increase the surface area of the electroplated metal composition.
  • the metal composition is electroplated on to a porous substrate.
  • the substrate comprises a planar surface and an edge or lip at the boundary of the planar surface and the cycles of differing electroplating conditions further inhibit an uneven deposition of the metal layer electrodeposited onto the planar surface and the edge or lip at the boundary of the planar surface.
  • the uneven deposition of the metal layer that is inhibited is a greater deposition of metal on the edge or lip at the boundary of the planar surface relative to the deposition of metal on the planar surface.
  • the electroplated metal composition comprises platinum.
  • the surface area of the third layer is at least 160, 170 or 180 times the geometric area of the third layer and is typically 230-260 times the geometric area of the third layer.
  • excessive metal e.g. platinum
  • growth may be observed along the periphery of the electrode made by electrodepositing layers of a metal(s) onto a substrate having a planar surface and the edge or lip at the boundary of the planar surface.
  • This “growth” is for example the result of an uneven deposition of the metal layer electrodeposited onto the planar surface and the edge or lip at the boundary of the planar surface.
  • using this current density under controlled conditions can result in a metal layer that is fragile and further has poor adhesion with the underlying substrate (e.g. a fragile platinum layer that exhibits a marginal adhesion to an underlying gold surface).
  • cyclic plating conditions are used that vary the applied current density during the plating process. Specific illustrative conditions are provided in Tables 1 and 2 below.
  • a low current density is first applied for 30 to 120 seconds (90 uA/sq. mm).
  • a higher current is then applied for 30 to 120 seconds (140 uA/sq. mm).
  • a low current density can be again applied for 30 to 120 seconds (90 uA/sq. mm).
  • Electrodes made by such specific combinations of cyclic plating conditions exhibit a number of surprising and beneficial material properties.
  • a first low current density cycle condition results in a smooth platinum layer, one that exhibits a strong adhesion to a substrate and a corresponding good resistance to abrasion (e.g. as measured by an art accepted abrasion test such as a DIN, TBS or Taber abrasion test method).
  • the first metal layer electroplated under a first set of conditions therefore exhibits a resistance to abrasion from the substrate that is greater than the resistance to abrasion exhibited by a metal layer electroplated to the substrate under the second set of conditions.
  • the cyclic plating conditions in this cycle therefore produce a metal layer that has the beneficial material property of good adherence between elements (e.g. the substrate and the first metal layer).
  • a second higher current density cycle is used to produce a further layer on this base layer on that exhibits a roughness that is correlated with a high surface area, a desirable property for electrodes that are for example coated with enzyme layer, such as an enzyme capable of reacting with and/or producing a molecule whose change in concentration can be measured by measuring a change in the current at the electrode (e.g. glucose oxidase).
  • electrodes made using these first and second conditions produce an electrodeposited metal layer where excessive “edge” growth (a factor that can lead to instability of electrode structures) is inhibited.
  • a further low density cycle plating cycle is then used to plate an additional metal layer that essentially hardens the rough layer of metal created in the second cycle (e.g. by making the layer more dense).
  • electrodes made by this cycle of plating conditions exhibit a constellation of desirable material properties including: (1) a good adhesion to the substrate (e.g. resulting from the conditions of cycle 1 ); (2) a desirable amount or degree of surface area (e.g. resulting from the conditions of cycle 2 ); and (3) a desirable hardness or density for final surface area of the electrode (e.g.
  • Electrodes plated in this fashion exhibit good adhesion to a gold substrate as well as good abrasion resistance. Furthermore, the resultant surface area ranges from 230-260 times the geometric surface area.
  • an implantable biosensor comprising an electrode comprising a plurality of electrodeposited metal layers including a first metal layer having a first surface area and a first adhesion strength with a substrate on which the first layer is electrodeposited and a second metal layer deposited on the first metal layer, the second metal layer having a second surface area and a second adhesion strength with the first layer on which the second layer is electrodeposited, wherein the second surface area is greater than the first surface area and the second adhesion strength is greater than the first adhesion strength; and further an enzyme layer disposed on the electrode, wherein an enzyme in the enzyme layer is capable of reacting with and/or producing a molecule whose change in concentration can be measured by measuring a change in the current at the electrode.
  • the electrode comprises a third metal layer deposited on the second metal layer, wherein the third metal layer has a greater density than the density of the second metal layer.
  • the enzyme layer comprises glucose oxidase or lactate oxidase.
  • Yet another embodiment of the invention is an electrode comprising a plurality of electrodeposited metal layers comprising a first metal layer having a first surface area; a second metal layer deposited on the first metal layer, the second metal layer having a second surface area that is greater than the first surface area of the first metal layer; and a third metal layer deposited on the second metal layer, the third metal layer having a greater density than the density of the second metal layer.
  • the substrate comprises a geometric feature selected to increase the surface area of the electrodeposited metal composition.
  • certain sensor embodiments achieve their bio-specificity through immobilized enzymes such as glucose oxidase (GOx) or lactate dehydrogenase (LDH), which consume oxygen along with glucose or lactate (the sensor analytes) as co-reactants.
  • GOx glucose oxidase
  • LDH lactate dehydrogenase
  • glucose or lactate the sensor analytes
  • a molar excess of oxygen is required.
  • glucose ( ⁇ 5 mM) and lactate ( ⁇ 1 mM) are almost always found in molar excess of oxygen ( ⁇ 0.05 mM).
  • most existing sensor designs employ a membrane that is significantly more permeable to oxygen than it is to the analyte.
  • permselective membranes usually contain poly(dimethylsiloxane) (PDMS), as it is biocompatible and typically possesses an unusually high permeability to oxygen and virtually no permeability to analytes such as glucose or lactate. Limited analyte permeability is typically imparted upon the PDMS-based material either by co-polymerizing PDMS with a hydrophilic polymer (i.e. Jeffamine®) or by cutting a macroscopic “window” into a tube or sheet of PDMS.
  • PDMS poly(dimethylsiloxane)
  • Copolymer-type permselective membranes have successfully been used in clinically approved short-term (less than 1-week) subcutaneous glucose sensors (e.g. Continuous Glucose Monitoring Systems (“CGMS) and/or a Telemetered Glucose Monitoring Systems (“TGMS”).
  • CGMS Continuous Glucose Monitoring Systems
  • TGMS Telemetered Glucose Monitoring Systems
  • This type of membrane simply requires analyte to diffuse across its thickness, which is optimal for sensor linearity and response time.
  • CGMS Continuous Glucose Monitoring Systems
  • TGMS Telemetered Glucose Monitoring Systems
  • Embodiments of the invention produce a permselective membrane fashioned from micro-fabricated PDMS that possesses the inherent advantages of both co-polymer and window-type biosensor membranes. While others have constructed PDMS microstructures through the casting of PDMS pre-polymers into complementary micro-fabricated relief patterns (e.g. Kumar et al., 1994 , Langmuir 10: 1498-1511; Dapprich, 2003, U.S. Pat. No. 6,585,939), no one has previously described the use of micro-fabricated PDMS as the permselective membrane in an enzymatic electrochemical biosensor.
  • photolithography, lithographic molding, thick-film printing, plasma polymerization (with or without shadow-masking), or discrete nano-dispensing can be used to micro-pattern a curable PDMS functionalized derivative, co-polymer, or mixture thereof onto a pre-cast immuno-isolation membrane.
  • Vacuum or a pressure gradient may or may not be applied to promote the filling of the pores of the immuno-isolation membrane.
  • Composite membranes fashioned in such a manner can possess morphologies that are layered, pore-filled, or some combination thereof.
  • a curable PDMS derivative, co-polymer, or mixture thereof can be micro-patterned onto a temporary substrate using the aforementioned techniques.
  • the stand-alone part may be used with a phase-inversion membrane (“PIM”) that may either be cast as a separate part or on top of the PDMS (filing its pores).
  • PIM phase-inversion membrane
  • a curable PDMS derivative, co-polymer, or mixture thereof can be micro-patterned directly onto the sensor assembly, using the techniques described above.
  • a PIM may be placed or cast on top of the micro-patterned PDMS.
  • Various methods for promoting adhesion may be employed by individuals skilled in the art.
  • a laser can be used to micro-machine holes (0.1-1000 microns) into a formed piece of PDMS co-polymer (or another polymeric composition such as “silicone rubber”) to form a microporous membrane.
  • the PDMS membrane may be used with or without a PIM in the final sensor assembly.
  • Various methods for promoting adhesion can be employed by individuals skilled in the art.
  • Illustrative chemically active groups that can be used to functionalize the PDMS and/or PDMS co-polymer include, but are not limited to: methacrylates, acrylates, vinyls, hydrides, silanols, alkoxys, amines, epoxides, carbinols, and mercaptos.
  • Examples of monomers that can be used to make the PDMS copolymer include, but are not limited to: phenylmethyl-, vinylmethyl-, diethyl-, methacryloxypropylmethyl-, acryloxypropylmethyl-, and alkylmethyl-siloxanes.
  • the immuno-isolation membrane can be pre-cast from a biocompatible polymer such as poly(acrylonitrile-vinyl chloride) (PAN-PVC), for example using a phase-inversion process that can be optimized by individuals well-trained in the fields of biomaterials and polymer chemistry.
  • PAN-PVC poly(acrylonitrile-vinyl chloride)
  • the casting of the phase-inversion membrane (PIM) and the micro-patterning of the PDMS can be performed on the sensor assembly itself or on a temporary substrate such as a glass slide or silicon wafer (e.g. to form a separate part).
  • a micro-patterned temporary substrate can also be used to create micro-wells in the PIM into which the PDMS could be patterned.
  • the permselective membranes containing poly(dimethylsiloxane) provide a variety of embodiments of the invention.
  • One embodiment of the invention is a method of making a membrane for use with an implantable analyte sensor by forming a first layer of material comprising a biocompatible polymer composition that is impermeable to immunoglobulins yet permeable to oxygen, glucose and lactate, and then coupling the first layer to a second layer comprising functionalized poly(dimethyl siloxane), functionalized poly(dimethyl siloxane) copolymer or a mixture of functionalized poly(dimethyl siloxane) and functionalized poly(dimethyl siloxane) copolymer so that a membrane for use with an implantable analyte is made.
  • a first layer of material comprising a biocompatible polymer composition that is impermeable to immunoglobulins is termed a “immuno-isolation membrane”.
  • the membrane made by this method is more permeable to oxygen than it is to compounds having a higher molecular weight such as glucose and/or lactate.
  • Composite membranes fashioned in such a manner can be made to possess a variety of morphologies, including those that are layered, pore-filled, or some combination thereof.
  • Illustrative chemically active groups that can be used to functionalize the PDMS and/or PDMS co-polymer include, but are not limited to: methacrylates, acrylates, vinyls, hydrides, silanols, alkoxys, amines, epoxides, carbinols, and mercaptos.
  • Examples of monomers that can be used to make the PDMS copolymer include, but are not limited to: phenylmethyl-, vinylmethyl-, diethyl-, methacryloxypropylmethyl-, acryloxypropylmethyl-, and alkylmethyl-siloxanes.
  • the other layer e.g.
  • the immuno-isolation membrane can be pre-cast from a biocompatible polymer such as poly(acrylonitrile-vinyl chloride) (PAN-PVC), for example using a phase-inversion process that can be optimized by individuals well-trained in the fields of biomaterials and polymer chemistry.
  • PAN-PVC poly(acrylonitrile-vinyl chloride)
  • the casting of the phase-inversion membrane (PIM) and the micro-patterning of the PDMS can be performed on the sensor assembly itself or on a temporary substrate such as a glass slide or silicon wafer (e.g. to form a separate part).
  • a micro-patterned temporary substrate can also be used to create micro-wells in the PIM into which the PDMS could be patterned. While certain embodiments of the invention include analyte sensors with composite membranes, the PDMS membrane may be used with or without a PIM in the final sensor assembly.
  • Embodiments of these membranes can be made using a variety of well-known techniques. For example, in one illustrative embodiment, photolithography, lithographic molding, thick-film printing, plasma polymerization (with or without shadow-masking), or discrete nano-dispensing can be used to micro-pattern a curable PDMS functionalized derivative, co-polymer, or mixture thereof onto a pre-cast immuno-isolation membrane. In another embodiment of the invention, a curable PDMS derivative, co-polymer, or mixture thereof can be micro-patterned onto a temporary substrate using the described techniques.
  • the stand-alone part may be used with a phase-inversion membrane that may either be cast as a separate part or on top of the PDMS (fling its pores).
  • a curable PDMS derivative, co-polymer, or mixture thereof can be micro-patterned directly onto the sensor assembly, using the techniques described above.
  • a PIM may be placed or cast on top of the micro-patterned PDMS.
  • an adhesive layer disposed between the first and second layers of the membrane to promote adhesion between the first and second layers (as well as any other sensor layer where such an adhesive layer is appropriate).
  • Various methods for promoting adhesion between the layers of the membrane may be employed by individuals skilled in the art.
  • a micro-patterned temporary substrate can also be used to create micro-wells in the PIM into which the PDMS could be patterned.
  • individuals skilled in the art can employ various chemicals and techniques for promoting adhesion between the PDMS and the PIM.
  • Examples include, but are not limited to the use of functionalized PDMS derivatives, silanes, silane esters, functionalized silane esters, cross-linking agents, reactive polymer coatings (see, e.g., Lahann et al., 2003 , Anal Chem. 75: 2117-2122), plasma treatment, plasma polymerization, shadow masking, and chemical vapor deposition.
  • the analyte sensor membrane can include additional layers having other compositions used in the manufacture of analyte sensors such as those described herein.
  • the first layer and/or the second layer of the membrane is constructed to include a plurality of pores.
  • a laser can be used to micro-machine holes (e.g. of about 0.1 to about 1000 microns in size) into a formed piece of PDMS co-polymer (or another polymeric composition such as “silicone rubber”) to form a microporous membrane.
  • At least one of the plurality of pores disposed in the second layer contains functionalized poly(dimethyl siloxane), functionalized poly(dimethyl siloxane) copolymer or a mixture of functionalized poly(dimethyl siloxane) and functionalized poly(dimethyl siloxane) copolymer of the first layer.
  • a related embodiment of the invention is a membrane made by the disclosed methods.
  • One such embodiment of the invention is a membrane for use with an implantable analyte sensor which includes a first layer comprising a biocompatible polymer composition that is impermeable to immunoglobulins, yet permeable to oxygen, glucose and lactate; and a second layer coupled to the first layer comprising functionalized poly(dimethyl siloxane), functionalized poly(dimethyl siloxane) copolymer or a mixture of functionalized poly(dimethyl siloxane) and functionalized poly(dimethyl siloxane) copolymer wherein the membrane is more permeable to oxygen than glucose and/or lactate.
  • the first and/or the second layers in the membrane comprises a plurality of pores.
  • at least one of the plurality of pores disposed in the second layer contains functionalized poly(dimethyl siloxane), functionalized poly(dimethyl siloxane) copolymer or a mixture of functionalized poly(dimethyl siloxane) and functionalized poly(dimethyl siloxane) copolymer of the second layer.
  • Yet another embodiment of the invention is an analyte sensor having a membrane disclosed above, for example an analyte sensor having a membrane made according the described methods.
  • a related embodiment is a method of making an analyte sensor having such a membrane.
  • Another embodiment of the invention is a membrane for use with an implantable analyte sensor, the membrane including a first layer comprising a biocompatible polymer composition that is impermeable to immunoglobulins, yet permeable to oxygen, glucose and lactate; and a second layer coupled to the first layer comprising functionalized poly(dimethyl siloxane), functionalized poly(dimethyl siloxane) copolymer or a mixture of functionalized poly(dimethyl siloxane) and functionalized poly(dimethyl siloxane) copolymer.
  • the membrane is typically more permeable to oxygen than glucose and/or lactate.
  • the first layer and/or the second layer comprises a plurality of pores disposed therein.
  • at least one of the plurality of pores disposed in the second layer contains functionalized poly(dimethyl siloxane), functionalized poly(dimethyl siloxane) copolymer or a mixture of functionalized poly(dimethyl siloxane) and functionalized poly(dimethyl siloxane) copolymer of the second layer.
  • an adhesive layer can be disposed between the first and second layers, wherein the adhesive layer promotes adhesion between the first and second layers.
  • a related embodiment of the invention is a method of making a membrane for use with an implantable analyte sensor generating a first layer comprising functionalized poly(dimethyl siloxane), functionalized poly(dimethyl siloxane) copolymer or a mixture of functionalized poly(dimethyl siloxane) and functionalized poly(dimethyl siloxane) copolymer; generating a second layer coupled to the first layer comprising a biocompatible polymer composition that is: impermeable to immunoglobulins; and permeable to oxygen, glucose and lactate so that a membrane is made that is more permeable to oxygen than glucose and/or lactate.
  • the first layer and/or the second layer can be made to comprise a plurality of pores disposed therein.
  • at least one of the plurality of pores disposed in the second layer is made to contain functionalized poly(dimethyl siloxane), functionalized poly(dimethyl siloxane) copolymer or a mixture of functionalized poly(dimethyl siloxane) and functionalized poly(dimethyl siloxane) copolymer of the second layer.
  • an adhesive layer can be disposed between the first and second layers.
  • the molds used in these procedures are usually fabricated using one of two different approaches.
  • the negative photoresist is coated, patterned via photolithography, and developed on a base substrate.
  • silicon wafers are etched to form a relief pattern.
  • the fabrication of molds with small, high aspect ratio features remains a significant challenge.
  • molds with these extreme geometries typically have poor mechanical properties and can for example detach from the underlying substrate during polymer release.
  • a base substrate formed from a material such as but not limited to glass, silicon, silicon nitride, or aluminum oxide is coated by a process such as sputtering with a conductive material such as but not limited to gold, silver, platinum, copper, or chrome.
  • the base substrate may be pre-coated with a layer of chrome or titanium in order to promote adhesion of the top conductive layer.
  • a positive photoresist such as AZ 4620 can then be applied by spin coating or by other methods familiar to those skilled in the art of microfabrication. After pre-baking, a sacrificial thin film of a metal such as chrome can be sputtered onto the layer of positive photo-resist.
  • a negative photoresist such as SU-8 can be applied by spin coating or by other methods familiar to those skilled in the art.
  • the layered substrate can be obtained through standard photolithography and subsequent development of the negative resist. Chrome etch or another appropriate etchant can be used to remove the areas of the sacrificial metal layer exposed by the development of the negative resist.
  • the substrate can then be re-exposed to UV light, with or without the use of a photomask.
  • the positive resist can then be developed using an appropriate developing solvent, which can be selected by individuals skilled in the art.
  • the resulting substrate can be electroplated with a conductive material such as, but not limited to gold, silver, platinum, copper, or chrome.
  • the positive resist, sacrificial metal layer, and the negative resist can be removed from the substrate by exposing it to acetone or another solvent that can be selected by those skilled in the art. If deemed necessary, other solvents and/or etchants such as chrome etch and negative resist strip can also be selected and applied by those skilled in the art.
  • the resulting mold can be used repeatedly to microfabricate PDMS and other elastomers.
  • the mold can also be used to microfabricate inelastic/hard materials in embodiments of the invention where the electroplated material can be removed by a chemical etchant or electrochemical oxidation.
  • One embodiment of the invention is a mold made by the methods described above.
  • a related embodiment of the invention is a mold for forming a polymerized composition having a predetermined geometry comprising a metallic substrate capable of containing a polymerizable composition, where the polymerized composition produced by the mold has an aspect ratio geometry selected to facilitate the selective diffusion or migration of a molecule involved in the sensing process.
  • Another embodiment of the invention is a mold for forming a polymerized composition having a predetermined geometry comprising a metallic substrate capable of containing a polymerizable composition, where the polymerized composition produced by the mold is between about 10 and 2000 microns in thickness.
  • Another embodiment of the invention is a mold for forming a polymerized composition having a predetermined geometry comprising a metallic substrate capable of containing a polymerizable composition, where the mold has sufficient mechanical strength to withstand release of a polymerized poly(dimethylsiloxane) composition without breaking.
  • the mold has a two or more of these features.
  • One embodiment of the invention is a method of making a mold for forming a polymerized composition of a predetermined geometry by providing a base substrate; disposing a conductive layer on to (at least a portion) the base substrate; disposing a positive photoresist layer on to the conductive layer; disposing a sacrificial metal layer on to the positive photoresist layer; disposing a negative photoresist layer on to the sacrificial metal layer; developing the negative photoresist layer via UV photolithography (with or without the use of a photomask); removing the areas of the sacrificial metal layer exposed by the development of the negative resist layer using an etchant; exposing these components to UV photolithography; developing the positive photoresist layer via a developing solvent; electroplating these components with a layer of conductive material; and removing the positive photoresist layer, the sacrificial metal layer, and the negative photoresist layer from the so layered substrate using a solvent so that the mold is made.
  • the base substrate can be formed from a wide variety of materials such as glass, silicon, silicon nitride, aluminum oxide or the like.
  • the conductive material is disposed on the base by a process such as sputtering. Conductive materials for use in embodiments of the invention include gold, silver, platinum, copper, chrome or the like.
  • the base substrate is coated with a layer of chrome or titanium prior to the application of the conductive material in order to promote adhesion of the base substrate and the conductive layer.
  • the substrate is baked prior to disposing the sacrificial metal layer on to the positive photoresist layer.
  • the negative photoresist layer and/or the positive photoresist layer is applied to the substrate by spin coating.
  • One embodiment of the invention is a mold for forming a polymerized composition having a predetermined geometry comprising a metallic substrate capable of containing a polymerizable composition; where the polymerized composition produced by the mold has an optimized aspect ratio geometry, the polymerized composition produced by the mold is between about 10 to 2000 microns in thickness; and the mold has sufficient mechanical strength to withstand release of a polymerized poly(dimethylsiloxane) composition without breaking.
  • a related embodiment of the invention is a method of making a mold for forming a polymerized composition of a predetermined geometry comprising: providing a base substrate; disposing a conductive layer on to (a portion) the base substrate; disposing a positive photoresist layer on to the conductive layer; disposing a sacrificial metal layer on to the positive photoresist layer; disposing a negative photoresist layer on to the sacrificial metal layer; developing the negative photoresist layer via UV photolithography; removing the areas of the sacrificial metal layer exposed by the development of the negative resist layer using an etchant; exposing these components to UV photolithography (with or without the use of a photomask); developing the positive photoresist layer via a developing solvent; electroplating these components with a layer of conductive material; and then removing the positive photoresist layer, the sacrificial metal layer, and the negative photoresist layer from the so layered substrate using a solvent so that the mold is made.
  • the mold made by the method can be used repeatedly.
  • the base substrate is coated with a layer of chrome or titanium prior to the application of the conductive layer in order to promote adhesion of the base substrate and the conductive layer.
  • the substrate is baked prior to disposing the sacrificial metal layer on to the positive photoresist layer.
  • the negative photoresist layer and/or the positive photoresist layer is applied to the substrate by spin coating.
  • a related embodiment of the invention is a method of sensing an analyte within the body of a mammal, the method comprising implanting an analyte sensor embodiment disclosed herein in to the mammal and then sensing an alteration in current at the working electrode and correlating the alteration in current with the presence of the analyte, so that the analyte is sensed.
  • the analyte sensor is polarized anodically such that the working electrode where the alteration in current is sensed is an anode.
  • the analyte sensor apparatus senses glucose in the mammal.
  • the analyte sensor apparatus senses lactate, potassium, calcium, oxygen, pH, and/or any physiologically relevant analyte in the mammal.
  • analyte sensors having the structure discussed above have a number of highly desirable characteristics which allow for a variety of methods for sensing analytes in a mammal.
  • the analyte sensor apparatus implanted in the mammal functions to sense an analyte within the body of a mammal for more than 1, 2, 3, 4, 5, or 6 months.
  • the analyte sensor apparatus so implanted in the mammal senses an alteration in current in response to an analyte within 15, 10, 5 or 2 minutes of the analyte contacting the sensor.
  • the sensors can be implanted into a variety of locations within the body of the mammal, for example in both vascular and non-vascular spaces.
  • kits and/or sensor set useful for the sensing an analyte as is described above.
  • the kit and/or sensor set typically comprises a container, a label and an analyte sensor as described above.
  • Suitable containers include, for example, an easy to open package made from a material such as a metal foil, bottles, vials, syringes, and test tubes.
  • the containers may be formed from a variety of materials such as metals (e.g. foils) paper products, glass or plastic.
  • the label on, or associated with, the container indicates that the sensor is used for assaying the analyte of choice.
  • the container holds a porous matrix that is coated with a layer of an enzyme such as glucose oxidase.
  • the kit and/or sensor set may further include other materials desirable from a commercial and user standpoint, including elements or devices designed to facilitate the introduction of the sensor into the analyte environment, other buffers, diluents, filters, needles, syringes, and package inserts with instructions for use.
  • Counter and ref Counter and ref Counter and ref Counter and ref Counter and ref electrodes get std electrodes get std electrodes get std electrodes get std plating for 2x3x plating for 2x3x plating for 2x3x plating for 2x3x plating for 2x3x config. config. config. config.

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US11/397,543 US20070227907A1 (en) 2006-04-04 2006-04-04 Methods and materials for controlling the electrochemistry of analyte sensors
US11/492,273 US9237865B2 (en) 2002-10-18 2006-07-25 Analyte sensors and methods for making and using them
US11/493,054 US9492111B2 (en) 2002-04-22 2006-07-26 Methods and materials for stabilizing analyte sensors
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Cited By (147)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US20080027287A1 (en) * 2002-04-22 2008-01-31 Rajiv Shah Methods and materials for stabilizing analyte sensors
US20090198117A1 (en) * 2008-01-29 2009-08-06 Medtronic Minimed, Inc. Analyte sensors having nanostructured electrodes and methods for making and using them
US20090227898A1 (en) * 2006-10-15 2009-09-10 Hans-Peter Haar Diagnostic test element and process for its production
US20090242525A1 (en) * 2008-03-31 2009-10-01 Electro Scientific Industries, Inc. Laser machining of fired ceramic and other hard and/or thick materials
US20090260448A1 (en) * 2008-04-18 2009-10-22 Korea Research Institute Of Standards And Science Apparatus for Measuring Pressure by Using Diaphragm and Method of Pressure Measurement by Using Diaphragm
US7613491B2 (en) 2002-05-22 2009-11-03 Dexcom, Inc. Silicone based membranes for use in implantable glucose sensors
US20090325221A1 (en) * 2008-06-30 2009-12-31 Kimberly-Clark Worldwide, Inc. Temporary Tattoo Decals for Detecting the Presence of an Analyte
US20100010323A1 (en) * 2008-03-21 2010-01-14 Nova Biomedical Corporation Temperature-compensated in-vivo sensor
US20100025238A1 (en) * 2008-07-31 2010-02-04 Medtronic Minimed, Inc. Analyte sensor apparatuses having improved electrode configurations and methods for making and using them
WO2010017119A1 (en) * 2008-08-06 2010-02-11 Reusense, Llc Reusable analyte sensing system and methods
WO2010027771A1 (en) * 2008-08-27 2010-03-11 Edwards Lifesciences Corporation Analyte sensor
US20100144047A1 (en) * 2008-12-04 2010-06-10 Waseda University Electrode for molecular sensing and method for molecular sensing
US20100159229A1 (en) * 2008-07-31 2010-06-24 Chae-Ho Shin Removal of bulge effects in nanopatterning
EP2206462A1 (en) * 2009-01-13 2010-07-14 Nederlandse Organisatie voor toegepast-natuurwetenschappelijk Onderzoek TNO A non-invasive chemical sensor, a skin patch, a packaging material and a monitoring system using the same
US7761130B2 (en) 2003-07-25 2010-07-20 Dexcom, Inc. Dual electrode system for a continuous analyte sensor
US20100200538A1 (en) * 2009-02-09 2010-08-12 Edwards Lifesciences Corporation Analyte Sensor and Fabrication Methods
US20100206746A1 (en) * 2007-05-30 2010-08-19 Bayer Healthcare, Llc Multi-layer pad and methods of using the same
US7792562B2 (en) 1997-03-04 2010-09-07 Dexcom, Inc. Device and method for determining analyte levels
US7828728B2 (en) 2003-07-25 2010-11-09 Dexcom, Inc. Analyte sensor
US20100285084A1 (en) * 2007-11-13 2010-11-11 Medtronic Minimed, Inc. Antimicrobial coatings for medical devices and methods for making and using them
US7875047B2 (en) 2002-04-19 2011-01-25 Pelikan Technologies, Inc. Method and apparatus for a multi-use body fluid sampling device with sterility barrier release
US7885697B2 (en) 2004-07-13 2011-02-08 Dexcom, Inc. Transcutaneous analyte sensor
US7882611B2 (en) * 2005-09-23 2011-02-08 Medtronic Minimed, Inc. Method of making an analyte sensor
US7892183B2 (en) 2002-04-19 2011-02-22 Pelikan Technologies, Inc. Method and apparatus for body fluid sampling and analyte sensing
US7901365B2 (en) 2002-04-19 2011-03-08 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US7909777B2 (en) 2002-04-19 2011-03-22 Pelikan Technologies, Inc Method and apparatus for penetrating tissue
US7909778B2 (en) 2002-04-19 2011-03-22 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US7909774B2 (en) 2002-04-19 2011-03-22 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US7909775B2 (en) 2001-06-12 2011-03-22 Pelikan Technologies, Inc. Method and apparatus for lancet launching device integrated onto a blood-sampling cartridge
US7914465B2 (en) 2002-04-19 2011-03-29 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US7976476B2 (en) 2002-04-19 2011-07-12 Pelikan Technologies, Inc. Device and method for variable speed lancet
US7981055B2 (en) 2001-06-12 2011-07-19 Pelikan Technologies, Inc. Tissue penetration device
US7981056B2 (en) 2002-04-19 2011-07-19 Pelikan Technologies, Inc. Methods and apparatus for lancet actuation
US7988645B2 (en) 2001-06-12 2011-08-02 Pelikan Technologies, Inc. Self optimizing lancing device with adaptation means to temporal variations in cutaneous properties
US20110196216A1 (en) * 2008-09-11 2011-08-11 Ortrud Quarder Electrode system for measuring an analyte concentration under in-vivo conditions
US8007446B2 (en) 2002-04-19 2011-08-30 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US8050731B2 (en) 2002-05-22 2011-11-01 Dexcom, Inc. Techniques to improve polyurethane membranes for implantable glucose sensors
US8062231B2 (en) 2002-04-19 2011-11-22 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
WO2011144904A1 (en) * 2010-05-19 2011-11-24 Lifescan Scotland Limited Analytical test strip with an electrode having electrochemically active and inert ares of a predetermined size and distribution
US20110301442A1 (en) * 2008-09-23 2011-12-08 Gilupi Gmbh Diagnostic analyte collection device based on flexible polymers with biological surface modification and microfluidic functionality
US8079960B2 (en) 2002-04-19 2011-12-20 Pelikan Technologies, Inc. Methods and apparatus for lancet actuation
WO2011115809A3 (en) * 2010-03-16 2012-01-19 Edwards Lifesciences Corporation High energy radiation insensitive analyte sensors
USRE43399E1 (en) 2003-07-25 2012-05-22 Dexcom, Inc. Electrode systems for electrochemical sensors
US8197421B2 (en) 2002-04-19 2012-06-12 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US8221334B2 (en) 2002-04-19 2012-07-17 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8255033B2 (en) 2003-07-25 2012-08-28 Dexcom, Inc. Oxygen enhancing membrane systems for implantable devices
US8251921B2 (en) 2003-06-06 2012-08-28 Sanofi-Aventis Deutschland Gmbh Method and apparatus for body fluid sampling and analyte sensing
US8262614B2 (en) 2003-05-30 2012-09-11 Pelikan Technologies, Inc. Method and apparatus for fluid injection
US8267870B2 (en) 2002-04-19 2012-09-18 Sanofi-Aventis Deutschland Gmbh Method and apparatus for body fluid sampling with hybrid actuation
US8277713B2 (en) 2004-05-03 2012-10-02 Dexcom, Inc. Implantable analyte sensor
US8282576B2 (en) 2003-09-29 2012-10-09 Sanofi-Aventis Deutschland Gmbh Method and apparatus for an improved sample capture device
US8287495B2 (en) 2009-07-30 2012-10-16 Tandem Diabetes Care, Inc. Infusion pump system with disposable cartridge having pressure venting and pressure feedback
US8296918B2 (en) 2003-12-31 2012-10-30 Sanofi-Aventis Deutschland Gmbh Method of manufacturing a fluid sampling device with improved analyte detecting member configuration
US8333710B2 (en) 2002-04-19 2012-12-18 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8360992B2 (en) 2002-04-19 2013-01-29 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8364229B2 (en) 2003-07-25 2013-01-29 Dexcom, Inc. Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise
US8372016B2 (en) 2002-04-19 2013-02-12 Sanofi-Aventis Deutschland Gmbh Method and apparatus for body fluid sampling and analyte sensing
US8382682B2 (en) 2002-04-19 2013-02-26 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8435190B2 (en) 2002-04-19 2013-05-07 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
WO2013066930A1 (en) 2011-11-01 2013-05-10 Medtronic Minimed, Inc. Methods and materials for modulating start-up time and air removal in dry sensors
US8439872B2 (en) 1998-03-30 2013-05-14 Sanofi-Aventis Deutschland Gmbh Apparatus and method for penetration with shaft having a sensor for sensing penetration depth
WO2013074668A1 (en) 2011-11-17 2013-05-23 Medtronic Minimed, Inc. Radition protecting composition and methods for making and using it
US8509871B2 (en) 2001-07-27 2013-08-13 Dexcom, Inc. Sensor head for use with implantable devices
US8560039B2 (en) 2008-09-19 2013-10-15 Dexcom, Inc. Particle-containing membrane and particulate electrode for analyte sensors
US8556829B2 (en) 2002-04-19 2013-10-15 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8562558B2 (en) 2007-06-08 2013-10-22 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US8573027B2 (en) 2009-02-27 2013-11-05 Tandem Diabetes Care, Inc. Methods and devices for determination of flow reservoir volume
US8574895B2 (en) 2002-12-30 2013-11-05 Sanofi-Aventis Deutschland Gmbh Method and apparatus using optical techniques to measure analyte levels
US8583204B2 (en) 2008-03-28 2013-11-12 Dexcom, Inc. Polymer membranes for continuous analyte sensors
WO2013177573A2 (en) 2012-05-25 2013-11-28 Medtronic Minimed, Inc. Foldover sensors and methods for making and using them
WO2014008297A1 (en) 2012-07-03 2014-01-09 Medtronic Minimed, Inc. Analyte sensors and production thereof
US8641644B2 (en) 2000-11-21 2014-02-04 Sanofi-Aventis Deutschland Gmbh Blood testing apparatus having a rotatable cartridge with multiple lancing elements and testing means
WO2014022641A1 (en) * 2012-08-01 2014-02-06 Massachusetts Institute Of Technology Actuation and control of stamp deformation in microcontact printing
US8652831B2 (en) 2004-12-30 2014-02-18 Sanofi-Aventis Deutschland Gmbh Method and apparatus for analyte measurement test time
US8668656B2 (en) 2003-12-31 2014-03-11 Sanofi-Aventis Deutschland Gmbh Method and apparatus for improving fluidic flow and sample capture
US8682408B2 (en) 2008-03-28 2014-03-25 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US8700114B2 (en) 2008-07-31 2014-04-15 Medtronic Minmed, Inc. Analyte sensor apparatuses comprising multiple implantable sensor elements and methods for making and using them
US8702624B2 (en) 2006-09-29 2014-04-22 Sanofi-Aventis Deutschland Gmbh Analyte measurement device with a single shot actuator
US8721671B2 (en) 2001-06-12 2014-05-13 Sanofi-Aventis Deutschland Gmbh Electric lancet actuator
US8744546B2 (en) 2005-05-05 2014-06-03 Dexcom, Inc. Cellulosic-based resistance domain for an analyte sensor
WO2014089276A1 (en) 2012-12-06 2014-06-12 Medtronic Minimed, Inc. Microarray electrodes useful with analyte sensors and methods for making and using them
US8784335B2 (en) 2002-04-19 2014-07-22 Sanofi-Aventis Deutschland Gmbh Body fluid sampling device with a capacitive sensor
WO2014116293A1 (en) 2013-01-22 2014-07-31 Medtronic Minimed, Inc. Muting glucose sensor oxygen response and reducing electrode edge growth with pulsed current plating
US8828201B2 (en) 2009-07-02 2014-09-09 Dexcom, Inc. Analyte sensors and methods of manufacturing same
US8828203B2 (en) 2004-05-20 2014-09-09 Sanofi-Aventis Deutschland Gmbh Printable hydrogels for biosensors
US8929968B2 (en) 2003-12-05 2015-01-06 Dexcom, Inc. Dual electrode system for a continuous analyte sensor
US8965476B2 (en) 2010-04-16 2015-02-24 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
WO2015069692A2 (en) 2013-11-07 2015-05-14 Medtronic Minimed, Inc. Enzyme matrices for use with ethylene oxide sterilization
US20150209733A1 (en) * 2012-06-26 2015-07-30 Fujifilm Manufacturing Europe Bv Process for Preparing Membranes
US9103727B2 (en) 2012-01-03 2015-08-11 The Regents Of The University Of Michigan Composite Fabry-Pérot sensor
US9144401B2 (en) 2003-06-11 2015-09-29 Sanofi-Aventis Deutschland Gmbh Low pain penetrating member
US9226699B2 (en) 2002-04-19 2016-01-05 Sanofi-Aventis Deutschland Gmbh Body fluid sampling module with a continuous compression tissue interface surface
US9248267B2 (en) 2002-04-19 2016-02-02 Sanofi-Aventis Deustchland Gmbh Tissue penetration device
US9250106B2 (en) 2009-02-27 2016-02-02 Tandem Diabetes Care, Inc. Methods and devices for determination of flow reservoir volume
US9314194B2 (en) 2002-04-19 2016-04-19 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US9324571B2 (en) * 2013-03-13 2016-04-26 Applied Materials, Inc. Post treatment for dielectric constant reduction with pore generation on low K dielectric films
US9351680B2 (en) 2003-10-14 2016-05-31 Sanofi-Aventis Deutschland Gmbh Method and apparatus for a variable user interface
US9351677B2 (en) 2009-07-02 2016-05-31 Dexcom, Inc. Analyte sensor with increased reference capacity
US9375169B2 (en) 2009-01-30 2016-06-28 Sanofi-Aventis Deutschland Gmbh Cam drive for managing disposable penetrating member actions with a single motor and motor and control system
US9386944B2 (en) 2008-04-11 2016-07-12 Sanofi-Aventis Deutschland Gmbh Method and apparatus for analyte detecting device
US9400241B2 (en) 2008-09-19 2016-07-26 Tandem Diabetes Care, Inc. Solute concentration measurement device and related methods
US9427532B2 (en) 2001-06-12 2016-08-30 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US20170108464A1 (en) * 2013-12-10 2017-04-20 Life Technologies Corporation System and method for forming microwells
US9763609B2 (en) 2003-07-25 2017-09-19 Dexcom, Inc. Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise
US9775553B2 (en) 2004-06-03 2017-10-03 Sanofi-Aventis Deutschland Gmbh Method and apparatus for a fluid sampling device
US9795747B2 (en) 2010-06-02 2017-10-24 Sanofi-Aventis Deutschland Gmbh Methods and apparatus for lancet actuation
US9820684B2 (en) 2004-06-03 2017-11-21 Sanofi-Aventis Deutschland Gmbh Method and apparatus for a fluid sampling device
EP3263712A1 (en) 2016-06-29 2018-01-03 Roche Diabetes Care GmbH Galvanically functionalized sensors
US9962486B2 (en) 2013-03-14 2018-05-08 Tandem Diabetes Care, Inc. System and method for detecting occlusions in an infusion pump
US9984891B2 (en) * 2016-04-21 2018-05-29 Shin-Etsu Chemical Co., Ltd. Method for forming organic film and method for manufacturing substrate for semiconductor apparatus
US9986942B2 (en) 2004-07-13 2018-06-05 Dexcom, Inc. Analyte sensor
WO2018170363A1 (en) 2017-03-17 2018-09-20 Medtronic Minimed, Inc. Metal pillar device structures and methods for making and using them in electrochemical and/or electrocatalytic applications
WO2019005687A1 (en) 2017-06-30 2019-01-03 Medtronic Minimed, Inc. SENSOR INITIALIZATION METHODS FOR FASTER BODY SENSOR RESPONSE
US10258736B2 (en) 2012-05-17 2019-04-16 Tandem Diabetes Care, Inc. Systems including vial adapter for fluid transfer
WO2019147578A1 (en) 2018-01-23 2019-08-01 Medtronic Minimed, Inc. Implantable polymer surfaces exhibiting reduced in vivo inflammatory responses
WO2019157106A2 (en) 2018-02-08 2019-08-15 Medtronic Minimed, Inc. Methods for controlling physical vapor deposition metal film adhesion to substrates and surfaces
WO2019157043A1 (en) 2018-02-08 2019-08-15 Medtronic Minimed, Inc. Glucose sensor electrode design
WO2019156934A1 (en) 2018-02-07 2019-08-15 Medtronic Minimed, Inc. Multilayer electrochemical analyte sensors and methods for making and using them
US10522398B2 (en) * 2017-08-31 2019-12-31 International Business Machines Corporation Modulating metal interconnect surface topography
US10610136B2 (en) 2005-03-10 2020-04-07 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
CN111742215A (zh) * 2018-04-09 2020-10-02 加州理工学院 金属-酶夹层
US10791928B2 (en) 2007-05-18 2020-10-06 Dexcom, Inc. Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise
US10813577B2 (en) 2005-06-21 2020-10-27 Dexcom, Inc. Analyte sensor
US10835672B2 (en) 2004-02-26 2020-11-17 Dexcom, Inc. Integrated insulin delivery system with continuous glucose sensor
EP3744248A1 (en) * 2019-05-29 2020-12-02 Bionime Corporation Biosensing device and activation method thereof
WO2021021538A1 (en) 2019-07-26 2021-02-04 Medtronic Minimed, Inc. Methods to improve oxygen delivery to implantable sensors
WO2021021867A1 (en) 2019-08-01 2021-02-04 Medtronic Minimed, Inc. Micro-pillar working electrodes design to reduce backflow of hydrogen peroxide in glucose sensor
US10966609B2 (en) 2004-02-26 2021-04-06 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
WO2022026542A1 (en) 2020-07-31 2022-02-03 Medtronic Minimed, Inc. Sensor identification and integrity check design
US11246990B2 (en) 2004-02-26 2022-02-15 Dexcom, Inc. Integrated delivery device for continuous glucose sensor
CN114224334A (zh) * 2021-12-28 2022-03-25 苏州原理科技有限公司 植入式生物传感器及制备方法
WO2022093574A1 (en) 2020-10-29 2022-05-05 Medtronic Minimed, Inc. Glucose biosensors comprising direct electron transfer enzymes and methods of making and using them
US11331022B2 (en) 2017-10-24 2022-05-17 Dexcom, Inc. Pre-connected analyte sensors
US11350862B2 (en) 2017-10-24 2022-06-07 Dexcom, Inc. Pre-connected analyte sensors
US11399745B2 (en) 2006-10-04 2022-08-02 Dexcom, Inc. Dual electrode system for a continuous analyte sensor
US11464908B2 (en) 2019-02-18 2022-10-11 Tandem Diabetes Care, Inc. Methods and apparatus for monitoring infusion sites for ambulatory infusion pumps
EP4071251A1 (en) 2021-04-09 2022-10-12 Medtronic MiniMed, Inc. Hexamethyldisiloxane membranes for analyte sensors
EP4134665A1 (en) 2021-08-13 2023-02-15 Medtronic MiniMed, Inc. Dry electrochemical impedance spectroscopy metrology for conductive chemical layers
CN115736909A (zh) * 2022-12-08 2023-03-07 深圳清华大学研究院 血糖检测装置及血糖检测系统
US11633133B2 (en) 2003-12-05 2023-04-25 Dexcom, Inc. Dual electrode system for a continuous analyte sensor
EP4174188A1 (en) 2021-10-14 2023-05-03 Medtronic Minimed, Inc. Sensors for 3-hydroxybutyrate detection
EP4190908A1 (en) 2021-12-02 2023-06-07 Medtronic Minimed, Inc. Ketone limiting membrane and dual layer membrane approach for ketone sensing
US11730407B2 (en) 2008-03-28 2023-08-22 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US20230277980A1 (en) * 2021-01-04 2023-09-07 Saudi Arabian Oil Company High free volume membrane for gas separation
EP4382611A1 (en) 2022-08-31 2024-06-12 Medtronic MiniMed, Inc. Sensors for 3-hydroxybutyrate detection
US20250194970A1 (en) * 2022-12-27 2025-06-19 Peking University Diabetes biosensor
US12517119B2 (en) 2021-04-19 2026-01-06 W. L. Gore & Associates, Inc. Electrochemical analyte biosensors, associated compositions of matter, and methods for making and using same

Families Citing this family (42)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US8527026B2 (en) 1997-03-04 2013-09-03 Dexcom, Inc. Device and method for determining analyte levels
US9282925B2 (en) 2002-02-12 2016-03-15 Dexcom, Inc. Systems and methods for replacing signal artifacts in a glucose sensor data stream
US8260393B2 (en) 2003-07-25 2012-09-04 Dexcom, Inc. Systems and methods for replacing signal data artifacts in a glucose sensor data stream
US8010174B2 (en) 2003-08-22 2011-08-30 Dexcom, Inc. Systems and methods for replacing signal artifacts in a glucose sensor data stream
US8282549B2 (en) 2003-12-09 2012-10-09 Dexcom, Inc. Signal processing for continuous analyte sensor
US7774145B2 (en) 2003-08-01 2010-08-10 Dexcom, Inc. Transcutaneous analyte sensor
US7519408B2 (en) 2003-11-19 2009-04-14 Dexcom, Inc. Integrated receiver for continuous analyte sensor
US8761856B2 (en) 2003-08-01 2014-06-24 Dexcom, Inc. System and methods for processing analyte sensor data
US20190357827A1 (en) 2003-08-01 2019-11-28 Dexcom, Inc. Analyte sensor
US20140121989A1 (en) 2003-08-22 2014-05-01 Dexcom, Inc. Systems and methods for processing analyte sensor data
US8364231B2 (en) 2006-10-04 2013-01-29 Dexcom, Inc. Analyte sensor
US8532730B2 (en) 2006-10-04 2013-09-10 Dexcom, Inc. Analyte sensor
US8452368B2 (en) 2004-07-13 2013-05-28 Dexcom, Inc. Transcutaneous analyte sensor
DE102008009817A1 (de) * 2008-02-19 2009-08-27 Epcos Ag Verbundwerkstoff zur Temperaturmessung, Temperatursensor aufweisend den Verbundwerkstoff und Verfahren zur Herstellung des Verbundwerkstoffs und des Temperatursensors
WO2009105709A1 (en) 2008-02-21 2009-08-27 Dexcom, Inc. Systems and methods for processing, transmitting and displaying sensor data
WO2010099507A1 (en) * 2009-02-26 2010-09-02 Abbott Diabetes Care Inc. Improved analyte sensors and methods of making and using the same
US20110004085A1 (en) 2009-04-30 2011-01-06 Dexcom, Inc. Performance reports associated with continuous sensor data from multiple analysis time periods
EP2482724A2 (en) 2009-09-30 2012-08-08 Dexcom, Inc. Transcutaneous analyte sensor
EP2545853B1 (en) 2010-03-09 2017-10-25 ARKRAY, Inc. Electrochemical sensor
US8608921B2 (en) * 2011-01-20 2013-12-17 Medtronic Minimed, Inc. Layered enzyme compositions for use with analyte sensors
EP3575796B1 (en) 2011-04-15 2020-11-11 DexCom, Inc. Advanced analyte sensor calibration and error detection
US9828620B2 (en) 2013-06-28 2017-11-28 Verily Life Sciences Llc Porous polymeric formulation prepared using monomer
US9750445B2 (en) 2013-06-28 2017-09-05 Verily Life Sciences Llc Porous polymeric formulation prepared using porogens
US20150160151A1 (en) * 2013-12-06 2015-06-11 Google Inc. Formulation and Storage Method to Enhance the Enzyme and Sensor Stabilities
TWI583947B (zh) * 2013-12-16 2017-05-21 聖高拜塑膠製品公司 電極及製造電極的方法
US9739746B1 (en) 2013-12-23 2017-08-22 Google Inc. Analyte sensor incorporating negatively charged moieties
US9834805B2 (en) 2013-12-23 2017-12-05 Verily Life Sciences Llc Two-layer analyte sensor
US9855359B2 (en) 2013-12-23 2018-01-02 Verily Life Sciences Llc Analyte sensors with ethylene oxide immunity
US9681828B2 (en) * 2014-05-01 2017-06-20 Medtronic Minimed, Inc. Physiological characteristic sensors and methods for forming such sensors
US10335302B2 (en) 2015-02-24 2019-07-02 Elira, Inc. Systems and methods for using transcutaneous electrical stimulation to enable dietary interventions
US20220062621A1 (en) 2015-02-24 2022-03-03 Elira, Inc. Electrical Stimulation-Based Weight Management System
US10864367B2 (en) 2015-02-24 2020-12-15 Elira, Inc. Methods for using an electrical dermal patch in a manner that reduces adverse patient reactions
US9956393B2 (en) 2015-02-24 2018-05-01 Elira, Inc. Systems for increasing a delay in the gastric emptying time for a patient using a transcutaneous electro-dermal patch
US10765863B2 (en) 2015-02-24 2020-09-08 Elira, Inc. Systems and methods for using a transcutaneous electrical stimulation device to deliver titrated therapy
US10376145B2 (en) 2015-02-24 2019-08-13 Elira, Inc. Systems and methods for enabling a patient to achieve a weight loss objective using an electrical dermal patch
CA2977584C (en) 2015-02-24 2024-03-05 Elira Therapeutics, Inc. Systems and methods for enabling appetite modulation and/or improving dietary compliance using an electro-dermal patch
US11298059B2 (en) 2016-05-13 2022-04-12 PercuSense, Inc. Analyte sensor
CN118533932A (zh) 2016-11-30 2024-08-23 美国圣戈班性能塑料公司 电极及电极制造方法
KR102050952B1 (ko) * 2018-02-20 2019-12-02 고려대학교 산학협력단 부착형 신축성 땀센서
TWI668444B (zh) * 2018-07-17 2019-08-11 淡江大學 檢測裝置及檢測系統
KR102699866B1 (ko) * 2021-12-29 2024-08-29 재단법인대구경북과학기술원 기계적 및 화학적 내구성이 향상된 생체전극 및 이의 제조방법
WO2025048947A1 (en) * 2023-08-31 2025-03-06 Virginia Tech Intellectual Properties, Inc. Assembly of flexible interconnects using sedimentation

Citations (24)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4971901A (en) * 1987-01-24 1990-11-20 Kanzaki Paper Manufacturing Co., Ltd. Process for preparing enzyme electrodes
US5063081A (en) * 1988-11-14 1991-11-05 I-Stat Corporation Method of manufacturing a plurality of uniform microfabricated sensing devices having an immobilized ligand receptor
US5310475A (en) * 1990-06-29 1994-05-10 Electroplating Engineers Of Japan, Limited Platinum electroforming and platinum electroplating
US5352348A (en) * 1987-04-09 1994-10-04 Nova Biomedical Corporation Method of using enzyme electrode
US5437999A (en) * 1994-02-22 1995-08-01 Boehringer Mannheim Corporation Electrochemical sensor
US5696314A (en) * 1996-07-12 1997-12-09 Chiron Diagnostics Corporation Multilayer enzyme electrode membranes and methods of making same
US5773270A (en) * 1991-03-12 1998-06-30 Chiron Diagnostics Corporation Three-layered membrane for use in an electrochemical sensor system
US5972199A (en) * 1995-10-11 1999-10-26 E. Heller & Company Electrochemical analyte sensors using thermostable peroxidase
US20010051768A1 (en) * 1997-10-20 2001-12-13 Joseph H Schulman Implantable enxzyme-based monitoring systems having improved longevity due to improved exterior surfaces
US6466810B1 (en) * 1995-11-22 2002-10-15 Legacy Good Samaritan Hospital And Medical Center Implantable device for monitoring changes in analyte concentration
US6484045B1 (en) * 2000-02-10 2002-11-19 Medtronic Minimed, Inc. Analyte sensor and method of making the same
US20030124623A1 (en) * 2001-12-05 2003-07-03 Paul Yager Microfluidic device and surface decoration process for solid phase affinity binding assays
US20040064156A1 (en) * 2002-09-27 2004-04-01 Medtronic Minimed, Inc. Method and apparatus for enhancing the integrity of an implantable sensor device
US20040074785A1 (en) * 2002-10-18 2004-04-22 Holker James D. Analyte sensors and methods for making them
US20040106857A1 (en) * 2001-02-15 2004-06-03 Regents Of The University Of California Membrane and electrode structure for implantable sensor
US20040194302A1 (en) * 1999-10-04 2004-10-07 Bhullar Raghbir S. Method of making a biosensor
US20050090607A1 (en) * 2003-10-28 2005-04-28 Dexcom, Inc. Silicone composition for biocompatible membrane
US20050101023A1 (en) * 2003-03-28 2005-05-12 Rogers James A. Methods for diagnosing urinary tract and prostatic disorders
US20050103625A1 (en) * 2001-07-27 2005-05-19 Rathbun Rhodes Sensor head for use with implantable devices
US20050205422A1 (en) * 2002-05-14 2005-09-22 Isabella Moser Enzyme electrode arrangement, a method for the fabrication of such arrangement and a biosensor arrangement comprising such electrode arrangement
US20050245799A1 (en) * 2004-05-03 2005-11-03 Dexcom, Inc. Implantable analyte sensor
US20050272989A1 (en) * 2004-06-04 2005-12-08 Medtronic Minimed, Inc. Analyte sensors and methods for making and using them
US20060052745A1 (en) * 2004-09-08 2006-03-09 Van Antwerp Nannette M Blood contacting sensor
US20070135698A1 (en) * 2005-12-13 2007-06-14 Rajiv Shah Biosensors and methods for making and using them

Family Cites Families (109)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US2899658A (en) * 1959-08-11 Leaf-type electrical resistance strain gage
US2402950A (en) * 1943-04-13 1946-07-02 Merlyn M Culver Molded part and method of forming same
US2519541A (en) * 1947-11-08 1950-08-22 Fluid Control Engineering Co Valve construction
US4034959A (en) * 1975-12-24 1977-07-12 Acf Industries, Incorporated Valve seal retaining means
US4104099A (en) * 1977-01-27 1978-08-01 International Telephone And Telegraph Corporation Method and apparatus for lapping or polishing materials
US4163544A (en) * 1977-11-10 1979-08-07 Acf Industries, Incorporated Two piece composite valve seal ring construction
US4373527B1 (en) * 1979-04-27 1995-06-27 Univ Johns Hopkins Implantable programmable medication infusion system
US4573994A (en) * 1979-04-27 1986-03-04 The Johns Hopkins University Refillable medication infusion apparatus
US4783441A (en) * 1979-04-30 1988-11-08 Hoechst Aktiengesellschaft Aqueous protein solutions stable to denaturation
US4356074A (en) * 1980-08-25 1982-10-26 The Yellow Springs Instrument Company, Inc. Substrate specific galactose oxidase enzyme electrodes
US4373009A (en) * 1981-05-18 1983-02-08 International Silicone Corporation Method of forming a hydrophilic coating on a substrate
US4569641A (en) * 1982-09-07 1986-02-11 Greatbatch Enterprises, Inc. Low power electromagnetic pump
US4714234A (en) * 1982-09-07 1987-12-22 Greatbatch Enterprises, Inc. Low power electromagnetic valve
US4568250A (en) * 1982-09-07 1986-02-04 Greatbatch Enterprises, Inc. Low power electromagnetic pump
US4769013A (en) * 1982-09-13 1988-09-06 Hydromer, Inc. Bio-effecting medical material and device
US4636150A (en) * 1983-05-23 1987-01-13 Greatbatch Enterprises, Inc. Low power electromagnetic pump
US4560375A (en) * 1983-06-30 1985-12-24 Pudenz-Schulte Medical Research Corp. Flow control valve
JPS60173452U (ja) 1984-04-27 1985-11-16 クラリオン株式会社 車両用盗難防止装置
IL73536A (en) * 1984-09-13 1987-12-20 Eltech Systems Corp Composite catalytic material particularly for electrolysis electrodes,its manufacture and its use in electrolysis
US4626244A (en) * 1985-02-01 1986-12-02 Consolidated Controls Corporation Implantable medication infusion device
DE3610316A1 (de) * 1985-03-29 1986-10-30 Four Brain K.K., Hiroshima Gebissbasis mit einer gummiartigen elastischen unterfutter- bzw. einlageschicht und verfahren zu seiner herstellung
WO1986007632A1 (en) * 1985-06-21 1986-12-31 Matsushita Electric Industrial Co., Ltd. Biosensor and method of manufacturing same
US4890620A (en) * 1985-09-20 1990-01-02 The Regents Of The University Of California Two-dimensional diffusion glucose substrate sensing electrode
US5019260A (en) * 1986-12-23 1991-05-28 Pall Corporation Filtration media with low protein adsorbability
US4759828A (en) * 1987-04-09 1988-07-26 Nova Biomedical Corporation Glucose electrode and method of determining glucose
US5001009A (en) * 1987-09-02 1991-03-19 Sterilization Technical Services, Inc. Lubricious hydrophilic composite coated on substrates
US4950256A (en) * 1988-04-07 1990-08-21 Luther Medical Products, Inc. Non-thrombogenic intravascular time release catheter
US5196088A (en) * 1988-08-05 1993-03-23 Tru Vue, Inc. Process and apparatus for producing non-glare glass by etching
US5212050A (en) * 1988-11-14 1993-05-18 Mier Randall M Method of forming a permselective layer
US5200051A (en) * 1988-11-14 1993-04-06 I-Stat Corporation Wholly microfabricated biosensors and process for the manufacture and use thereof
US5077210A (en) * 1989-01-13 1991-12-31 Eigler Frances S Immobilization of active agents on substrates with a silane and heterobifunctional crosslinking agent
EP0397130B1 (en) * 1989-05-11 1995-04-19 Kanegafuchi Kagaku Kogyo Kabushiki Kaisha Medical device having highly biocompatible surface and method for manufacturing the same
US4958661A (en) * 1989-08-08 1990-09-25 The Lee Company Check valve
US5330911A (en) * 1989-09-28 1994-07-19 Board Of Regents, The University Of Texas System Surfaces having desirable cell adhesive effects
US5288636A (en) * 1989-12-15 1994-02-22 Boehringer Mannheim Corporation Enzyme electrode system
DE58904449D1 (de) * 1989-12-21 1993-06-24 Siemens Ag Anordnung zum transkutanen ein- oder nachfuellen von fluessigen medikamenten in ein implantierbares medikamentendosiergeraet.
US5108819A (en) * 1990-02-14 1992-04-28 Eli Lilly And Company Thin film electrical component
US5165407A (en) * 1990-04-19 1992-11-24 The University Of Kansas Implantable glucose sensor
DE4018597A1 (de) * 1990-06-09 1991-12-12 Draegerwerk Ag Messelektrode fuer eine elektrochemische gasmesszelle
JP3017816B2 (ja) * 1991-02-20 2000-03-13 清原 まさ子 流体制御器
US5593852A (en) * 1993-12-02 1997-01-14 Heller; Adam Subcutaneous glucose electrode
WO1993001221A1 (en) * 1991-07-05 1993-01-21 Biocompatibles Limited Polymeric surface coatings
US5322063A (en) * 1991-10-04 1994-06-21 Eli Lilly And Company Hydrophilic polyurethane membranes for electrochemical glucose sensors
ZA938555B (en) * 1992-11-23 1994-08-02 Lilly Co Eli Technique to improve the performance of electrochemical sensors
US5308641A (en) * 1993-01-19 1994-05-03 Medtronic, Inc. Biocompatibility of solid surfaces
US5229172A (en) * 1993-01-19 1993-07-20 Medtronic, Inc. Modification of polymeric surface by graft polymerization
US5299571A (en) * 1993-01-22 1994-04-05 Eli Lilly And Company Apparatus and method for implantation of sensors
US5433710A (en) * 1993-03-16 1995-07-18 Minimed, Inc. Medication infusion pump with fluoropolymer valve seat
US5328954A (en) * 1993-04-16 1994-07-12 Icet, Inc. Encrusting and bacterial resistant coatings for medical applications
CA2174246C (en) * 1993-10-29 2004-05-11 Gary R. Schildgen Method of manufacturing a medical electrical lead
US5390691A (en) * 1994-01-27 1995-02-21 Sproule; Ronald Bleed valve for water supply for camping vehicle
US5391250A (en) * 1994-03-15 1995-02-21 Minimed Inc. Method of fabricating thin film sensors
US5390671A (en) * 1994-03-15 1995-02-21 Minimed Inc. Transcutaneous sensor insertion set
ATE202942T1 (de) * 1994-04-15 2001-07-15 Cobe Cardiovascular Inc Blutverträglicher beschichteter artikel
US5482473A (en) * 1994-05-09 1996-01-09 Minimed Inc. Flex circuit connector
US5662960A (en) * 1995-02-01 1997-09-02 Schneider (Usa) Inc. Process for producing slippery, tenaciously adhering hydrogel coatings containing a polyurethane-urea polymer hydrogel commingled with a poly (n-vinylpyrrolidone) polymer hydrogel
US5568806A (en) * 1995-02-16 1996-10-29 Minimed Inc. Transcutaneous sensor insertion set
US5702754A (en) * 1995-02-22 1997-12-30 Meadox Medicals, Inc. Method of providing a substrate with a hydrophilic coating and substrates, particularly medical devices, provided with such coatings
JPH08247988A (ja) * 1995-03-07 1996-09-27 Nippon Telegr & Teleph Corp <Ntt> 電気化学検出器及びその製造方法
AU5363596A (en) * 1995-03-17 1996-10-08 Regents Of The University Of Minnesota Biocompatible materials
US5882494A (en) * 1995-03-27 1999-03-16 Minimed, Inc. Polyurethane/polyurea compositions containing silicone for biosensor membranes
US5786439A (en) * 1996-10-24 1998-07-28 Minimed Inc. Hydrophilic, swellable coatings for biosensors
US5609629A (en) * 1995-06-07 1997-03-11 Med Institute, Inc. Coated implantable medical device
US5620738A (en) * 1995-06-07 1997-04-15 Union Carbide Chemicals & Plastics Technology Corporation Non-reactive lubicious coating process
US5672638A (en) * 1995-08-22 1997-09-30 Medtronic, Inc. Biocompatability for solid surfaces
US5607475A (en) * 1995-08-22 1997-03-04 Medtronic, Inc. Biocompatible medical article and method
US5913040A (en) * 1995-08-22 1999-06-15 Backweb Ltd. Method and apparatus for transmitting and displaying information between a remote network and a local computer
US5716506A (en) * 1995-10-06 1998-02-10 Board Of Trustees Of The University Of Illinois Electrochemical sensors for gas detection
US5809242A (en) * 1996-04-19 1998-09-15 Juno Online Services, L.P. Electronic mail system for displaying advertisement at local computer received from remote system while the local computer is off-line the remote system
US5728420A (en) 1996-08-09 1998-03-17 Medtronic, Inc. Oxidative method for attachment of glycoproteins to surfaces of medical devices
US6013855A (en) * 1996-08-06 2000-01-11 United States Surgical Grafting of biocompatible hydrophilic polymers onto inorganic and metal surfaces
US5804048A (en) * 1996-08-15 1998-09-08 Via Medical Corporation Electrode assembly for assaying glucose
KR0176334B1 (ko) 1996-08-19 1999-04-01 박원훈 내인체 감염성 삽입금속표면의 코팅방법 및 그 치료기술
WO1998008553A1 (en) 1996-08-29 1998-03-05 Medtronic, Inc. Biocompatibility for solid surfaces
US5855618A (en) 1996-09-13 1999-01-05 Meadox Medicals, Inc. Polyurethanes grafted with polyethylene oxide chains containing covalently bonded heparin
DE69733307T2 (de) 1996-11-01 2006-01-19 Medtronic MiniMed, Inc., Northridge Pumpe zur medikamentösen infusion mit einer durch ein protein stabilisierten beschichtung
JP2943739B2 (ja) * 1996-12-13 1999-08-30 日本電気株式会社 尿糖計および尿糖測定方法
US5964993A (en) * 1996-12-19 1999-10-12 Implanted Biosystems Inc. Glucose sensor
US5858186A (en) * 1996-12-20 1999-01-12 The Regents Of The University Of California Urea biosensor for hemodialysis monitoring
CA2278709A1 (en) * 1997-01-27 1998-08-13 Benjamin Slotznick System for delivering and displaying primary and secondary information
US6558351B1 (en) * 1999-06-03 2003-05-06 Medtronic Minimed, Inc. Closed loop system for controlling insulin infusion
US5954643A (en) 1997-06-09 1999-09-21 Minimid Inc. Insertion set for a transcutaneous sensor
US6411998B1 (en) * 1997-09-08 2002-06-25 International Business Machines Corporation World wide web internet delay monitor
US5965244A (en) 1997-10-24 1999-10-12 Rexam Graphics Inc. Printing medium comprised of porous medium
US5880364A (en) 1997-10-31 1999-03-09 Cosense, Inc. Non-contact ultrasonic micromeasurement system
US5939208A (en) * 1997-11-07 1999-08-17 Biomimetics, Inc. Method for creation of biomimetic surfaces
US6579690B1 (en) * 1997-12-05 2003-06-17 Therasense, Inc. Blood analyte monitoring through subcutaneous measurement
US6103033A (en) 1998-03-04 2000-08-15 Therasense, Inc. Process for producing an electrochemical biosensor
US6134461A (en) 1998-03-04 2000-10-17 E. Heller & Company Electrochemical analyte
US6175752B1 (en) 1998-04-30 2001-01-16 Therasense, Inc. Analyte monitoring device and methods of use
DE19932749B4 (de) * 1998-07-23 2006-05-04 Robert Bosch Gmbh Schichtsystem und Verfahren zu dessen Herstellung sowie dessen Verwendung
US6554798B1 (en) * 1998-08-18 2003-04-29 Medtronic Minimed, Inc. External infusion device with remote programming, bolus estimator and/or vibration alarm capabilities
ATE269114T1 (de) * 1998-11-20 2004-07-15 Univ Connecticut Verfahren und vorrichtung zur steuerung der gewebeimplantat-interaktionen
US6368274B1 (en) * 1999-07-01 2002-04-09 Medtronic Minimed, Inc. Reusable analyte sensor site and method of using the same
US6413393B1 (en) * 1999-07-07 2002-07-02 Minimed, Inc. Sensor including UV-absorbing polymer and method of manufacture
US6645359B1 (en) * 2000-10-06 2003-11-11 Roche Diagnostics Corporation Biosensor
US20020139668A1 (en) * 1999-11-03 2002-10-03 Raghbir Singh Bhullar Embedded metallic deposits
EP1253852B1 (en) 2000-02-10 2007-09-19 Medtronic MiniMed, Inc. Analyte sensor
US6340421B1 (en) * 2000-05-16 2002-01-22 Minimed Inc. Microelectrogravimetric method for plating a biosensor
US6642015B2 (en) 2000-12-29 2003-11-04 Minimed Inc. Hydrophilic polymeric material for coating biosensors
EP1345057A3 (de) * 2002-03-16 2004-08-04 Nexans Verfahren zur Herstellung eines optischen Kabels
EP1418428A1 (en) * 2002-11-07 2004-05-12 GenOdyssee Method to provide natural therapeutic agents with high therapeutic index
US20070088208A1 (en) * 2003-02-17 2007-04-19 Mikito Yasuzawa Linear device
US7718439B2 (en) * 2003-06-20 2010-05-18 Roche Diagnostics Operations, Inc. System and method for coding information on a biosensor test strip
JP2005060166A (ja) * 2003-08-12 2005-03-10 Canon Inc カーボン被覆構造体およびその製造方法、及びカーボンチューブの製造方法、及びカーボン被覆電極、及び機能性素子
US7387714B2 (en) * 2003-11-06 2008-06-17 3M Innovative Properties Company Electrochemical sensor strip
US8277713B2 (en) * 2004-05-03 2012-10-02 Dexcom, Inc. Implantable analyte sensor
US20070045902A1 (en) 2004-07-13 2007-03-01 Brauker James H Analyte sensor
WO2006026748A1 (en) * 2004-08-31 2006-03-09 Lifescan Scotland Limited Method of manufacturing an auto-calibrating sensor

Patent Citations (26)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4971901A (en) * 1987-01-24 1990-11-20 Kanzaki Paper Manufacturing Co., Ltd. Process for preparing enzyme electrodes
US5352348A (en) * 1987-04-09 1994-10-04 Nova Biomedical Corporation Method of using enzyme electrode
US5063081A (en) * 1988-11-14 1991-11-05 I-Stat Corporation Method of manufacturing a plurality of uniform microfabricated sensing devices having an immobilized ligand receptor
US5310475A (en) * 1990-06-29 1994-05-10 Electroplating Engineers Of Japan, Limited Platinum electroforming and platinum electroplating
US5773270A (en) * 1991-03-12 1998-06-30 Chiron Diagnostics Corporation Three-layered membrane for use in an electrochemical sensor system
US5437999A (en) * 1994-02-22 1995-08-01 Boehringer Mannheim Corporation Electrochemical sensor
US5972199A (en) * 1995-10-11 1999-10-26 E. Heller & Company Electrochemical analyte sensors using thermostable peroxidase
US6466810B1 (en) * 1995-11-22 2002-10-15 Legacy Good Samaritan Hospital And Medical Center Implantable device for monitoring changes in analyte concentration
US5696314A (en) * 1996-07-12 1997-12-09 Chiron Diagnostics Corporation Multilayer enzyme electrode membranes and methods of making same
US20010051768A1 (en) * 1997-10-20 2001-12-13 Joseph H Schulman Implantable enxzyme-based monitoring systems having improved longevity due to improved exterior surfaces
US20040194302A1 (en) * 1999-10-04 2004-10-07 Bhullar Raghbir S. Method of making a biosensor
US6484045B1 (en) * 2000-02-10 2002-11-19 Medtronic Minimed, Inc. Analyte sensor and method of making the same
US20040106857A1 (en) * 2001-02-15 2004-06-03 Regents Of The University Of California Membrane and electrode structure for implantable sensor
US20050103625A1 (en) * 2001-07-27 2005-05-19 Rathbun Rhodes Sensor head for use with implantable devices
US20030124623A1 (en) * 2001-12-05 2003-07-03 Paul Yager Microfluidic device and surface decoration process for solid phase affinity binding assays
US20050205422A1 (en) * 2002-05-14 2005-09-22 Isabella Moser Enzyme electrode arrangement, a method for the fabrication of such arrangement and a biosensor arrangement comprising such electrode arrangement
US20040064156A1 (en) * 2002-09-27 2004-04-01 Medtronic Minimed, Inc. Method and apparatus for enhancing the integrity of an implantable sensor device
US20040074785A1 (en) * 2002-10-18 2004-04-22 Holker James D. Analyte sensors and methods for making them
US20050101023A1 (en) * 2003-03-28 2005-05-12 Rogers James A. Methods for diagnosing urinary tract and prostatic disorders
US20050090607A1 (en) * 2003-10-28 2005-04-28 Dexcom, Inc. Silicone composition for biocompatible membrane
US20050245799A1 (en) * 2004-05-03 2005-11-03 Dexcom, Inc. Implantable analyte sensor
US20050272989A1 (en) * 2004-06-04 2005-12-08 Medtronic Minimed, Inc. Analyte sensors and methods for making and using them
US20060052745A1 (en) * 2004-09-08 2006-03-09 Van Antwerp Nannette M Blood contacting sensor
US20060293576A1 (en) * 2004-09-08 2006-12-28 Van Antwerp Nanette M Protrudent analyte sensor
US7468033B2 (en) * 2004-09-08 2008-12-23 Medtronic Minimed, Inc. Blood contacting sensor
US20070135698A1 (en) * 2005-12-13 2007-06-14 Rajiv Shah Biosensors and methods for making and using them

Cited By (370)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US8527025B1 (en) 1997-03-04 2013-09-03 Dexcom, Inc. Device and method for determining analyte levels
US7974672B2 (en) 1997-03-04 2011-07-05 Dexcom, Inc. Device and method for determining analyte levels
US7970448B2 (en) 1997-03-04 2011-06-28 Dexcom, Inc. Device and method for determining analyte levels
US7835777B2 (en) 1997-03-04 2010-11-16 Dexcom, Inc. Device and method for determining analyte levels
US7792562B2 (en) 1997-03-04 2010-09-07 Dexcom, Inc. Device and method for determining analyte levels
US8676288B2 (en) 1997-03-04 2014-03-18 Dexcom, Inc. Device and method for determining analyte levels
US9339223B2 (en) 1997-03-04 2016-05-17 Dexcom, Inc. Device and method for determining analyte levels
US8439872B2 (en) 1998-03-30 2013-05-14 Sanofi-Aventis Deutschland Gmbh Apparatus and method for penetration with shaft having a sensor for sensing penetration depth
US8641644B2 (en) 2000-11-21 2014-02-04 Sanofi-Aventis Deutschland Gmbh Blood testing apparatus having a rotatable cartridge with multiple lancing elements and testing means
US8206317B2 (en) 2001-06-12 2012-06-26 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8282577B2 (en) 2001-06-12 2012-10-09 Sanofi-Aventis Deutschland Gmbh Method and apparatus for lancet launching device integrated onto a blood-sampling cartridge
US8845550B2 (en) 2001-06-12 2014-09-30 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US7981055B2 (en) 2001-06-12 2011-07-19 Pelikan Technologies, Inc. Tissue penetration device
US7988645B2 (en) 2001-06-12 2011-08-02 Pelikan Technologies, Inc. Self optimizing lancing device with adaptation means to temporal variations in cutaneous properties
US7909775B2 (en) 2001-06-12 2011-03-22 Pelikan Technologies, Inc. Method and apparatus for lancet launching device integrated onto a blood-sampling cartridge
US8641643B2 (en) 2001-06-12 2014-02-04 Sanofi-Aventis Deutschland Gmbh Sampling module device and method
US9802007B2 (en) 2001-06-12 2017-10-31 Sanofi-Aventis Deutschland Gmbh Methods and apparatus for lancet actuation
US8382683B2 (en) 2001-06-12 2013-02-26 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8360991B2 (en) 2001-06-12 2013-01-29 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US9694144B2 (en) 2001-06-12 2017-07-04 Sanofi-Aventis Deutschland Gmbh Sampling module device and method
US8343075B2 (en) 2001-06-12 2013-01-01 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8337421B2 (en) 2001-06-12 2012-12-25 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8622930B2 (en) 2001-06-12 2014-01-07 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8679033B2 (en) 2001-06-12 2014-03-25 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8216154B2 (en) 2001-06-12 2012-07-10 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8211037B2 (en) 2001-06-12 2012-07-03 Pelikan Technologies, Inc. Tissue penetration device
US8206319B2 (en) 2001-06-12 2012-06-26 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8721671B2 (en) 2001-06-12 2014-05-13 Sanofi-Aventis Deutschland Gmbh Electric lancet actuator
US8162853B2 (en) 2001-06-12 2012-04-24 Pelikan Technologies, Inc. Tissue penetration device
US8123700B2 (en) 2001-06-12 2012-02-28 Pelikan Technologies, Inc. Method and apparatus for lancet launching device integrated onto a blood-sampling cartridge
US9427532B2 (en) 2001-06-12 2016-08-30 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8016774B2 (en) 2001-06-12 2011-09-13 Pelikan Technologies, Inc. Tissue penetration device
US9937298B2 (en) 2001-06-12 2018-04-10 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8509871B2 (en) 2001-07-27 2013-08-13 Dexcom, Inc. Sensor head for use with implantable devices
US9804114B2 (en) 2001-07-27 2017-10-31 Dexcom, Inc. Sensor head for use with implantable devices
US9328371B2 (en) 2001-07-27 2016-05-03 Dexcom, Inc. Sensor head for use with implantable devices
US9560993B2 (en) 2001-11-21 2017-02-07 Sanofi-Aventis Deutschland Gmbh Blood testing apparatus having a rotatable cartridge with multiple lancing elements and testing means
US8556829B2 (en) 2002-04-19 2013-10-15 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8372016B2 (en) 2002-04-19 2013-02-12 Sanofi-Aventis Deutschland Gmbh Method and apparatus for body fluid sampling and analyte sensing
US7976476B2 (en) 2002-04-19 2011-07-12 Pelikan Technologies, Inc. Device and method for variable speed lancet
US7938787B2 (en) 2002-04-19 2011-05-10 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US7981056B2 (en) 2002-04-19 2011-07-19 Pelikan Technologies, Inc. Methods and apparatus for lancet actuation
US7988644B2 (en) 2002-04-19 2011-08-02 Pelikan Technologies, Inc. Method and apparatus for a multi-use body fluid sampling device with sterility barrier release
US7914465B2 (en) 2002-04-19 2011-03-29 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US7909774B2 (en) 2002-04-19 2011-03-22 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US8007446B2 (en) 2002-04-19 2011-08-30 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US7909778B2 (en) 2002-04-19 2011-03-22 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US8845549B2 (en) 2002-04-19 2014-09-30 Sanofi-Aventis Deutschland Gmbh Method for penetrating tissue
US8808201B2 (en) 2002-04-19 2014-08-19 Sanofi-Aventis Deutschland Gmbh Methods and apparatus for penetrating tissue
US7909777B2 (en) 2002-04-19 2011-03-22 Pelikan Technologies, Inc Method and apparatus for penetrating tissue
US8062231B2 (en) 2002-04-19 2011-11-22 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US8905945B2 (en) 2002-04-19 2014-12-09 Dominique M. Freeman Method and apparatus for penetrating tissue
US8784335B2 (en) 2002-04-19 2014-07-22 Sanofi-Aventis Deutschland Gmbh Body fluid sampling device with a capacitive sensor
US9072842B2 (en) 2002-04-19 2015-07-07 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8079960B2 (en) 2002-04-19 2011-12-20 Pelikan Technologies, Inc. Methods and apparatus for lancet actuation
US9724021B2 (en) 2002-04-19 2017-08-08 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US9089294B2 (en) 2002-04-19 2015-07-28 Sanofi-Aventis Deutschland Gmbh Analyte measurement device with a single shot actuator
US7901365B2 (en) 2002-04-19 2011-03-08 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US9089678B2 (en) 2002-04-19 2015-07-28 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8157748B2 (en) 2002-04-19 2012-04-17 Pelikan Technologies, Inc. Methods and apparatus for lancet actuation
US7892183B2 (en) 2002-04-19 2011-02-22 Pelikan Technologies, Inc. Method and apparatus for body fluid sampling and analyte sensing
US8690796B2 (en) 2002-04-19 2014-04-08 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8197421B2 (en) 2002-04-19 2012-06-12 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US8197423B2 (en) 2002-04-19 2012-06-12 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US8202231B2 (en) 2002-04-19 2012-06-19 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US9795334B2 (en) 2002-04-19 2017-10-24 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US9498160B2 (en) 2002-04-19 2016-11-22 Sanofi-Aventis Deutschland Gmbh Method for penetrating tissue
US7875047B2 (en) 2002-04-19 2011-01-25 Pelikan Technologies, Inc. Method and apparatus for a multi-use body fluid sampling device with sterility barrier release
US9186468B2 (en) 2002-04-19 2015-11-17 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8221334B2 (en) 2002-04-19 2012-07-17 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8235915B2 (en) 2002-04-19 2012-08-07 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US9226699B2 (en) 2002-04-19 2016-01-05 Sanofi-Aventis Deutschland Gmbh Body fluid sampling module with a continuous compression tissue interface surface
US8636673B2 (en) 2002-04-19 2014-01-28 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US9248267B2 (en) 2002-04-19 2016-02-02 Sanofi-Aventis Deustchland Gmbh Tissue penetration device
US8579831B2 (en) 2002-04-19 2013-11-12 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8574168B2 (en) 2002-04-19 2013-11-05 Sanofi-Aventis Deutschland Gmbh Method and apparatus for a multi-use body fluid sampling device with analyte sensing
US8267870B2 (en) 2002-04-19 2012-09-18 Sanofi-Aventis Deutschland Gmbh Method and apparatus for body fluid sampling with hybrid actuation
US8562545B2 (en) 2002-04-19 2013-10-22 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US9314194B2 (en) 2002-04-19 2016-04-19 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US9907502B2 (en) 2002-04-19 2018-03-06 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8496601B2 (en) 2002-04-19 2013-07-30 Sanofi-Aventis Deutschland Gmbh Methods and apparatus for lancet actuation
US8491500B2 (en) 2002-04-19 2013-07-23 Sanofi-Aventis Deutschland Gmbh Methods and apparatus for lancet actuation
US9339612B2 (en) 2002-04-19 2016-05-17 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8333710B2 (en) 2002-04-19 2012-12-18 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8435190B2 (en) 2002-04-19 2013-05-07 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8337419B2 (en) 2002-04-19 2012-12-25 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US8337420B2 (en) 2002-04-19 2012-12-25 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
US9839386B2 (en) 2002-04-19 2017-12-12 Sanofi-Aventis Deustschland Gmbh Body fluid sampling device with capacitive sensor
US8430828B2 (en) 2002-04-19 2013-04-30 Sanofi-Aventis Deutschland Gmbh Method and apparatus for a multi-use body fluid sampling device with sterility barrier release
US8360992B2 (en) 2002-04-19 2013-01-29 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8414503B2 (en) 2002-04-19 2013-04-09 Sanofi-Aventis Deutschland Gmbh Methods and apparatus for lancet actuation
US8366637B2 (en) 2002-04-19 2013-02-05 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US7959582B2 (en) 2002-04-19 2011-06-14 Pelikan Technologies, Inc. Method and apparatus for penetrating tissue
US8403864B2 (en) 2002-04-19 2013-03-26 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8382682B2 (en) 2002-04-19 2013-02-26 Sanofi-Aventis Deutschland Gmbh Method and apparatus for penetrating tissue
US8388551B2 (en) 2002-04-19 2013-03-05 Sanofi-Aventis Deutschland Gmbh Method and apparatus for multi-use body fluid sampling device with sterility barrier release
US9492111B2 (en) 2002-04-22 2016-11-15 Medtronic Minimed, Inc. Methods and materials for stabilizing analyte sensors
US20080027287A1 (en) * 2002-04-22 2008-01-31 Rajiv Shah Methods and materials for stabilizing analyte sensors
US10154807B2 (en) 2002-05-22 2018-12-18 Dexcom, Inc. Techniques to improve polyurethane membranes for implantable glucose sensors
US9549693B2 (en) 2002-05-22 2017-01-24 Dexcom, Inc. Silicone based membranes for use in implantable glucose sensors
US8064977B2 (en) 2002-05-22 2011-11-22 Dexcom, Inc. Silicone based membranes for use in implantable glucose sensors
US8050731B2 (en) 2002-05-22 2011-11-01 Dexcom, Inc. Techniques to improve polyurethane membranes for implantable glucose sensors
US10052051B2 (en) 2002-05-22 2018-08-21 Dexcom, Inc. Silicone based membranes for use in implantable glucose sensors
US11020026B2 (en) 2002-05-22 2021-06-01 Dexcom, Inc. Silicone based membranes for use in implantable glucose sensors
US8865249B2 (en) 2002-05-22 2014-10-21 Dexcom, Inc. Techniques to improve polyurethane membranes for implantable glucose sensors
US8053018B2 (en) 2002-05-22 2011-11-08 Dexcom, Inc. Techniques to improve polyurethane membranes for implantable glucose sensors
US9179869B2 (en) 2002-05-22 2015-11-10 Dexcom, Inc. Techniques to improve polyurethane membranes for implantable glucose sensors
US7613491B2 (en) 2002-05-22 2009-11-03 Dexcom, Inc. Silicone based membranes for use in implantable glucose sensors
US8543184B2 (en) 2002-05-22 2013-09-24 Dexcom, Inc. Silicone based membranes for use in implantable glucose sensors
US9801574B2 (en) 2002-05-22 2017-10-31 Dexcom, Inc. Techniques to improve polyurethane membranes for implantable glucose sensors
US8574895B2 (en) 2002-12-30 2013-11-05 Sanofi-Aventis Deutschland Gmbh Method and apparatus using optical techniques to measure analyte levels
US9034639B2 (en) 2002-12-30 2015-05-19 Sanofi-Aventis Deutschland Gmbh Method and apparatus using optical techniques to measure analyte levels
US8262614B2 (en) 2003-05-30 2012-09-11 Pelikan Technologies, Inc. Method and apparatus for fluid injection
US8251921B2 (en) 2003-06-06 2012-08-28 Sanofi-Aventis Deutschland Gmbh Method and apparatus for body fluid sampling and analyte sensing
US9144401B2 (en) 2003-06-11 2015-09-29 Sanofi-Aventis Deutschland Gmbh Low pain penetrating member
US10034628B2 (en) 2003-06-11 2018-07-31 Sanofi-Aventis Deutschland Gmbh Low pain penetrating member
US7828728B2 (en) 2003-07-25 2010-11-09 Dexcom, Inc. Analyte sensor
US8364229B2 (en) 2003-07-25 2013-01-29 Dexcom, Inc. Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise
US8909314B2 (en) 2003-07-25 2014-12-09 Dexcom, Inc. Oxygen enhancing membrane systems for implantable devices
US10376143B2 (en) 2003-07-25 2019-08-13 Dexcom, Inc. Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise
US7761130B2 (en) 2003-07-25 2010-07-20 Dexcom, Inc. Dual electrode system for a continuous analyte sensor
US9763609B2 (en) 2003-07-25 2017-09-19 Dexcom, Inc. Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise
US8255030B2 (en) 2003-07-25 2012-08-28 Dexcom, Inc. Oxygen enhancing membrane systems for implantable devices
US10610140B2 (en) 2003-07-25 2020-04-07 Dexcom, Inc. Oxygen enhancing membrane systems for implantable devices
US9597027B2 (en) 2003-07-25 2017-03-21 Dexcom, Inc. Oxygen enhancing membrane systems for implantable devices
US9993186B2 (en) 2003-07-25 2018-06-12 Dexcom, Inc. Oxygen enhancing membrane systems for implantable devices
US8255032B2 (en) 2003-07-25 2012-08-28 Dexcom, Inc. Oxygen enhancing membrane systems for implantable devices
USRE43399E1 (en) 2003-07-25 2012-05-22 Dexcom, Inc. Electrode systems for electrochemical sensors
US8255033B2 (en) 2003-07-25 2012-08-28 Dexcom, Inc. Oxygen enhancing membrane systems for implantable devices
US8945910B2 (en) 2003-09-29 2015-02-03 Sanofi-Aventis Deutschland Gmbh Method and apparatus for an improved sample capture device
US8282576B2 (en) 2003-09-29 2012-10-09 Sanofi-Aventis Deutschland Gmbh Method and apparatus for an improved sample capture device
US9351680B2 (en) 2003-10-14 2016-05-31 Sanofi-Aventis Deutschland Gmbh Method and apparatus for a variable user interface
US10188333B2 (en) 2003-12-05 2019-01-29 Dexcom, Inc. Calibration techniques for a continuous analyte sensor
US8929968B2 (en) 2003-12-05 2015-01-06 Dexcom, Inc. Dual electrode system for a continuous analyte sensor
US11633133B2 (en) 2003-12-05 2023-04-25 Dexcom, Inc. Dual electrode system for a continuous analyte sensor
US8668656B2 (en) 2003-12-31 2014-03-11 Sanofi-Aventis Deutschland Gmbh Method and apparatus for improving fluidic flow and sample capture
US8296918B2 (en) 2003-12-31 2012-10-30 Sanofi-Aventis Deutschland Gmbh Method of manufacturing a fluid sampling device with improved analyte detecting member configuration
US9561000B2 (en) 2003-12-31 2017-02-07 Sanofi-Aventis Deutschland Gmbh Method and apparatus for improving fluidic flow and sample capture
US10835672B2 (en) 2004-02-26 2020-11-17 Dexcom, Inc. Integrated insulin delivery system with continuous glucose sensor
US12102410B2 (en) 2004-02-26 2024-10-01 Dexcom, Inc Integrated medicament delivery device for use with continuous analyte sensor
US11246990B2 (en) 2004-02-26 2022-02-15 Dexcom, Inc. Integrated delivery device for continuous glucose sensor
US10966609B2 (en) 2004-02-26 2021-04-06 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US12226617B2 (en) 2004-02-26 2025-02-18 Dexcom, Inc. Integrated delivery device for continuous glucose sensor
US12115357B2 (en) 2004-02-26 2024-10-15 Dexcom, Inc. Integrated delivery device for continuous glucose sensor
US8277713B2 (en) 2004-05-03 2012-10-02 Dexcom, Inc. Implantable analyte sensor
US9261476B2 (en) 2004-05-20 2016-02-16 Sanofi Sa Printable hydrogel for biosensors
US8828203B2 (en) 2004-05-20 2014-09-09 Sanofi-Aventis Deutschland Gmbh Printable hydrogels for biosensors
US9820684B2 (en) 2004-06-03 2017-11-21 Sanofi-Aventis Deutschland Gmbh Method and apparatus for a fluid sampling device
US9775553B2 (en) 2004-06-03 2017-10-03 Sanofi-Aventis Deutschland Gmbh Method and apparatus for a fluid sampling device
US11045120B2 (en) 2004-07-13 2021-06-29 Dexcom, Inc. Analyte sensor
US10524703B2 (en) 2004-07-13 2020-01-07 Dexcom, Inc. Transcutaneous analyte sensor
US11883164B2 (en) 2004-07-13 2024-01-30 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10799158B2 (en) 2004-07-13 2020-10-13 Dexcom, Inc. Analyte sensor
US10799159B2 (en) 2004-07-13 2020-10-13 Dexcom, Inc. Analyte sensor
US7885697B2 (en) 2004-07-13 2011-02-08 Dexcom, Inc. Transcutaneous analyte sensor
US9986942B2 (en) 2004-07-13 2018-06-05 Dexcom, Inc. Analyte sensor
US8792953B2 (en) 2004-07-13 2014-07-29 Dexcom, Inc. Transcutaneous analyte sensor
US10993641B2 (en) 2004-07-13 2021-05-04 Dexcom, Inc. Analyte sensor
US10722152B2 (en) 2004-07-13 2020-07-28 Dexcom, Inc. Analyte sensor
US10813576B2 (en) 2004-07-13 2020-10-27 Dexcom, Inc. Analyte sensor
US10932700B2 (en) 2004-07-13 2021-03-02 Dexcom, Inc. Analyte sensor
US10993642B2 (en) 2004-07-13 2021-05-04 Dexcom, Inc. Analyte sensor
US10709362B2 (en) 2004-07-13 2020-07-14 Dexcom, Inc. Analyte sensor
US10709363B2 (en) 2004-07-13 2020-07-14 Dexcom, Inc. Analyte sensor
US10827956B2 (en) 2004-07-13 2020-11-10 Dexcom, Inc. Analyte sensor
US9414777B2 (en) 2004-07-13 2016-08-16 Dexcom, Inc. Transcutaneous analyte sensor
US10918314B2 (en) 2004-07-13 2021-02-16 Dexcom, Inc. Analyte sensor
US11064917B2 (en) 2004-07-13 2021-07-20 Dexcom, Inc. Analyte sensor
US10918315B2 (en) 2004-07-13 2021-02-16 Dexcom, Inc. Analyte sensor
US11026605B1 (en) 2004-07-13 2021-06-08 Dexcom, Inc. Analyte sensor
US10980452B2 (en) 2004-07-13 2021-04-20 Dexcom, Inc. Analyte sensor
US10918313B2 (en) 2004-07-13 2021-02-16 Dexcom, Inc. Analyte sensor
US8652831B2 (en) 2004-12-30 2014-02-18 Sanofi-Aventis Deutschland Gmbh Method and apparatus for analyte measurement test time
US10617336B2 (en) 2005-03-10 2020-04-14 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10925524B2 (en) 2005-03-10 2021-02-23 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10918316B2 (en) 2005-03-10 2021-02-16 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10610136B2 (en) 2005-03-10 2020-04-07 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10898114B2 (en) 2005-03-10 2021-01-26 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10918318B2 (en) 2005-03-10 2021-02-16 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US11051726B2 (en) 2005-03-10 2021-07-06 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10856787B2 (en) 2005-03-10 2020-12-08 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10610137B2 (en) 2005-03-10 2020-04-07 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10918317B2 (en) 2005-03-10 2021-02-16 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10610135B2 (en) 2005-03-10 2020-04-07 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US11000213B2 (en) 2005-03-10 2021-05-11 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10709364B2 (en) 2005-03-10 2020-07-14 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10716498B2 (en) 2005-03-10 2020-07-21 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10743801B2 (en) 2005-03-10 2020-08-18 Dexcom, Inc. System and methods for processing analyte sensor data for sensor calibration
US10300507B2 (en) 2005-05-05 2019-05-28 Dexcom, Inc. Cellulosic-based resistance domain for an analyte sensor
US8744546B2 (en) 2005-05-05 2014-06-03 Dexcom, Inc. Cellulosic-based resistance domain for an analyte sensor
US10813577B2 (en) 2005-06-21 2020-10-27 Dexcom, Inc. Analyte sensor
US9693722B2 (en) 2005-09-23 2017-07-04 Medtronic Minimed, Inc. Sensor with layered electrodes
US7882611B2 (en) * 2005-09-23 2011-02-08 Medtronic Minimed, Inc. Method of making an analyte sensor
US8850687B2 (en) 2005-09-23 2014-10-07 Medtronic Minimed, Inc. Method of making a sensor
US9179870B2 (en) 2005-09-23 2015-11-10 Medtronic Minimed, Inc. Method of making a sensor with layered electrodes
US9364177B2 (en) 2005-09-23 2016-06-14 Medtronic Minimed, Inc. Sensor with layered electrodes
US8850688B2 (en) 2005-09-23 2014-10-07 Medtronic Minimed, Inc. Method of making a sensor with layered electrodes
US8702624B2 (en) 2006-09-29 2014-04-22 Sanofi-Aventis Deutschland Gmbh Analyte measurement device with a single shot actuator
US11399745B2 (en) 2006-10-04 2022-08-02 Dexcom, Inc. Dual electrode system for a continuous analyte sensor
US8052618B2 (en) * 2006-10-15 2011-11-08 Roche Diagnostics Operations, Inc. Diagnostic test element and process for its production
US20090227898A1 (en) * 2006-10-15 2009-09-10 Hans-Peter Haar Diagnostic test element and process for its production
US10791928B2 (en) 2007-05-18 2020-10-06 Dexcom, Inc. Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise
US12433485B2 (en) 2007-05-18 2025-10-07 Dexcom, Inc. Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise
US9993202B2 (en) * 2007-05-30 2018-06-12 Ascensia Diabetes Care Holdings Ag Multi-layer pad and methods of using the same
US20100206746A1 (en) * 2007-05-30 2010-08-19 Bayer Healthcare, Llc Multi-layer pad and methods of using the same
US12394120B2 (en) 2007-06-08 2025-08-19 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US10403012B2 (en) 2007-06-08 2019-09-03 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US12536716B2 (en) 2007-06-08 2026-01-27 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US9741139B2 (en) 2007-06-08 2017-08-22 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US11373347B2 (en) 2007-06-08 2022-06-28 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US8562558B2 (en) 2007-06-08 2013-10-22 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US12456239B2 (en) 2007-06-08 2025-10-28 Dexcom, Inc. Integrated medicament delivery device for use with continuous analyte sensor
US12246166B2 (en) 2007-10-09 2025-03-11 Dexcom, Inc. Integrated insulin delivery system with continuous glucose sensor
US11160926B1 (en) 2007-10-09 2021-11-02 Dexcom, Inc. Pre-connected analyte sensors
US12397113B2 (en) 2007-10-09 2025-08-26 Dexcom, Inc. Integrated insulin delivery system with continuous glucose sensor
US12397110B2 (en) 2007-10-09 2025-08-26 Dexcom, Inc. Integrated insulin delivery system with continuous glucose sensor
US11744943B2 (en) 2007-10-09 2023-09-05 Dexcom, Inc. Integrated insulin delivery system with continuous glucose sensor
US8512731B2 (en) 2007-11-13 2013-08-20 Medtronic Minimed, Inc. Antimicrobial coatings for medical devices and methods for making and using them
US20100285084A1 (en) * 2007-11-13 2010-11-11 Medtronic Minimed, Inc. Antimicrobial coatings for medical devices and methods for making and using them
US9309550B2 (en) 2008-01-29 2016-04-12 Medtronic Minimed, Inc. Analyte sensors having nanostructured electrodes and methods for making and using them
US20090198117A1 (en) * 2008-01-29 2009-08-06 Medtronic Minimed, Inc. Analyte sensors having nanostructured electrodes and methods for making and using them
US20100010323A1 (en) * 2008-03-21 2010-01-14 Nova Biomedical Corporation Temperature-compensated in-vivo sensor
US9173607B2 (en) 2008-03-28 2015-11-03 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US8583204B2 (en) 2008-03-28 2013-11-12 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US11147483B2 (en) 2008-03-28 2021-10-19 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US9572523B2 (en) 2008-03-28 2017-02-21 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US9566026B2 (en) 2008-03-28 2017-02-14 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US9549699B2 (en) 2008-03-28 2017-01-24 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US9693721B2 (en) 2008-03-28 2017-07-04 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US8954128B2 (en) 2008-03-28 2015-02-10 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US8682408B2 (en) 2008-03-28 2014-03-25 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US11730407B2 (en) 2008-03-28 2023-08-22 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US9173606B2 (en) 2008-03-28 2015-11-03 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US10143410B2 (en) 2008-03-28 2018-12-04 Dexcom, Inc. Polymer membranes for continuous analyte sensors
US20090242525A1 (en) * 2008-03-31 2009-10-01 Electro Scientific Industries, Inc. Laser machining of fired ceramic and other hard and/or thick materials
US8723075B2 (en) 2008-03-31 2014-05-13 Electro Scientific Industries, Inc. Laser machining of fired ceramic and other hard and/or thick materials
US8093532B2 (en) * 2008-03-31 2012-01-10 Electro Scientific Industries, Inc. Laser machining of fired ceramic and other hard and/or thick materials
US9386944B2 (en) 2008-04-11 2016-07-12 Sanofi-Aventis Deutschland Gmbh Method and apparatus for analyte detecting device
US7784352B2 (en) * 2008-04-18 2010-08-31 Korea Research Institute Of Standards And Science Apparatus for measuring pressure by using diaphragm and method of pressure measurement by using diaphragm
US20090260448A1 (en) * 2008-04-18 2009-10-22 Korea Research Institute Of Standards And Science Apparatus for Measuring Pressure by Using Diaphragm and Method of Pressure Measurement by Using Diaphragm
WO2010001275A3 (en) * 2008-06-30 2010-03-25 Kimberly-Clark Worldwide, Inc. Temporary tattoo decals for detecting the presence of an analyte
US20090325221A1 (en) * 2008-06-30 2009-12-31 Kimberly-Clark Worldwide, Inc. Temporary Tattoo Decals for Detecting the Presence of an Analyte
US20100159229A1 (en) * 2008-07-31 2010-06-24 Chae-Ho Shin Removal of bulge effects in nanopatterning
US20100025238A1 (en) * 2008-07-31 2010-02-04 Medtronic Minimed, Inc. Analyte sensor apparatuses having improved electrode configurations and methods for making and using them
US8700114B2 (en) 2008-07-31 2014-04-15 Medtronic Minmed, Inc. Analyte sensor apparatuses comprising multiple implantable sensor elements and methods for making and using them
US10327678B2 (en) 2008-07-31 2019-06-25 Medtronic Minimed, Inc. Analyte sensor apparatuses comprising multiple implantable sensor elements and methods for making and using them
WO2010017119A1 (en) * 2008-08-06 2010-02-11 Reusense, Llc Reusable analyte sensing system and methods
US20100243477A1 (en) * 2008-08-27 2010-09-30 Edwards Lifesciences Corporation Analyte Sensor
US8900431B2 (en) 2008-08-27 2014-12-02 Edwards Lifesciences Corporation Analyte sensor
WO2010027771A1 (en) * 2008-08-27 2010-03-11 Edwards Lifesciences Corporation Analyte sensor
US10390743B2 (en) * 2008-09-11 2019-08-27 Roche Diabetes Care, Inc. Electrode system for measuring an analyte concentration under in-vivo conditions
US9394563B2 (en) * 2008-09-11 2016-07-19 Roche Diabetes Care, Inc. Electrode system for measuring an analyte concentration under in-vivo conditions
US20110196216A1 (en) * 2008-09-11 2011-08-11 Ortrud Quarder Electrode system for measuring an analyte concentration under in-vivo conditions
US10028683B2 (en) 2008-09-19 2018-07-24 Dexcom, Inc. Particle-containing membrane and particulate electrode for analyte sensors
US9400241B2 (en) 2008-09-19 2016-07-26 Tandem Diabetes Care, Inc. Solute concentration measurement device and related methods
US8560039B2 (en) 2008-09-19 2013-10-15 Dexcom, Inc. Particle-containing membrane and particulate electrode for analyte sensors
US9339222B2 (en) 2008-09-19 2016-05-17 Dexcom, Inc. Particle-containing membrane and particulate electrode for analyte sensors
US10028684B2 (en) 2008-09-19 2018-07-24 Dexcom, Inc. Particle-containing membrane and particulate electrode for analyte sensors
US10561352B2 (en) 2008-09-19 2020-02-18 Dexcom, Inc. Particle-containing membrane and particulate electrode for analyte sensors
US11918354B2 (en) 2008-09-19 2024-03-05 Dexcom, Inc. Particle-containing membrane and particulate electrode for analyte sensors
US20110301442A1 (en) * 2008-09-23 2011-12-08 Gilupi Gmbh Diagnostic analyte collection device based on flexible polymers with biological surface modification and microfluidic functionality
US10314524B2 (en) * 2008-09-23 2019-06-11 Gilupi Gmbh Diagnostic analyte collection device based on flexible polymers with biological surface modification and microfluidic functionality
US20100144047A1 (en) * 2008-12-04 2010-06-10 Waseda University Electrode for molecular sensing and method for molecular sensing
US8148160B2 (en) * 2008-12-04 2012-04-03 Waseda University Electrode for molecular sensing and method for molecular sensing
WO2010082820A3 (en) * 2009-01-13 2010-12-02 Nederlandse Organisatie Voor Toegepast-Natuurwetenschappelijk Onderzoek Tno A non-invasive chemical sensor, a skin patch, a packaging material and a monitoring system using the same
EP2206462A1 (en) * 2009-01-13 2010-07-14 Nederlandse Organisatie voor toegepast-natuurwetenschappelijk Onderzoek TNO A non-invasive chemical sensor, a skin patch, a packaging material and a monitoring system using the same
US9375169B2 (en) 2009-01-30 2016-06-28 Sanofi-Aventis Deutschland Gmbh Cam drive for managing disposable penetrating member actions with a single motor and motor and control system
US8696917B2 (en) * 2009-02-09 2014-04-15 Edwards Lifesciences Corporation Analyte sensor and fabrication methods
US20140166612A1 (en) * 2009-02-09 2014-06-19 Edwards Lifesciences Corporation Analyte sensor and fabrication methods
US20100200538A1 (en) * 2009-02-09 2010-08-12 Edwards Lifesciences Corporation Analyte Sensor and Fabrication Methods
US8573027B2 (en) 2009-02-27 2013-11-05 Tandem Diabetes Care, Inc. Methods and devices for determination of flow reservoir volume
US9250106B2 (en) 2009-02-27 2016-02-02 Tandem Diabetes Care, Inc. Methods and devices for determination of flow reservoir volume
US9763608B2 (en) 2009-07-02 2017-09-19 Dexcom, Inc. Analyte sensors and methods of manufacturing same
US8828201B2 (en) 2009-07-02 2014-09-09 Dexcom, Inc. Analyte sensors and methods of manufacturing same
US9237864B2 (en) 2009-07-02 2016-01-19 Dexcom, Inc. Analyte sensors and methods of manufacturing same
US9517025B2 (en) 2009-07-02 2016-12-13 Dexcom, Inc. Analyte sensor with increased reference capacity
US12220234B2 (en) 2009-07-02 2025-02-11 Dexcom, Inc. Analyte sensors and methods of manufacturing same
US9351677B2 (en) 2009-07-02 2016-05-31 Dexcom, Inc. Analyte sensor with increased reference capacity
US10470691B2 (en) 2009-07-02 2019-11-12 Dexcom, Inc. Analyte sensor with increased reference capacity
US12409515B2 (en) 2009-07-02 2025-09-09 Dexcom, Inc. Continuous analyte sensors and methods of making same
US9131885B2 (en) 2009-07-02 2015-09-15 Dexcom, Inc. Analyte sensors and methods of manufacturing same
US8926561B2 (en) 2009-07-30 2015-01-06 Tandem Diabetes Care, Inc. Infusion pump system with disposable cartridge having pressure venting and pressure feedback
US12042627B2 (en) 2009-07-30 2024-07-23 Tandem Diabetes Care, Inc. Infusion pump systems and methods
US8758323B2 (en) 2009-07-30 2014-06-24 Tandem Diabetes Care, Inc. Infusion pump system with disposable cartridge having pressure venting and pressure feedback
US12144964B2 (en) 2009-07-30 2024-11-19 Tandem Diabetes Care, Inc Infusion pump system with disposable cartridge having pressure venting and pressure feedback
US11135362B2 (en) 2009-07-30 2021-10-05 Tandem Diabetes Care, Inc. Infusion pump systems and methods
US11285263B2 (en) 2009-07-30 2022-03-29 Tandem Diabetes Care, Inc. Infusion pump systems and methods
US12539364B2 (en) 2009-07-30 2026-02-03 Tandem Diabetes Care, Inc. Infusion pump systems with disposable cartridge having pressure venting and pressure feedback
US9211377B2 (en) 2009-07-30 2015-12-15 Tandem Diabetes Care, Inc. Infusion pump system with disposable cartridge having pressure venting and pressure feedback
US8298184B2 (en) 2009-07-30 2012-10-30 Tandem Diabetes Care, Inc. Infusion pump system with disposable cartridge having pressure venting and pressure feedback
US8287495B2 (en) 2009-07-30 2012-10-16 Tandem Diabetes Care, Inc. Infusion pump system with disposable cartridge having pressure venting and pressure feedback
WO2011115809A3 (en) * 2010-03-16 2012-01-19 Edwards Lifesciences Corporation High energy radiation insensitive analyte sensors
US8965476B2 (en) 2010-04-16 2015-02-24 Sanofi-Aventis Deutschland Gmbh Tissue penetration device
WO2011144904A1 (en) * 2010-05-19 2011-11-24 Lifescan Scotland Limited Analytical test strip with an electrode having electrochemically active and inert ares of a predetermined size and distribution
US8932449B2 (en) 2010-05-19 2015-01-13 Lifescan Scotland Limited Analytical test strip with an electrode having electrochemically active and inert areas of a predetermined size and distribution
US8940141B2 (en) 2010-05-19 2015-01-27 Lifescan Scotland Limited Analytical test strip with an electrode having electrochemically active and inert areas of a predetermined size and distribution
EP2647992A1 (en) * 2010-05-19 2013-10-09 Lifescan Scotland Limited Analytical test strip with an electrode having electrochemically active and inert areas of a predetermined size and distribution
US9795747B2 (en) 2010-06-02 2017-10-24 Sanofi-Aventis Deutschland Gmbh Methods and apparatus for lancet actuation
US9989522B2 (en) 2011-11-01 2018-06-05 Medtronic Minimed, Inc. Methods and materials for modulating start-up time and air removal in dry sensors
WO2013066930A1 (en) 2011-11-01 2013-05-10 Medtronic Minimed, Inc. Methods and materials for modulating start-up time and air removal in dry sensors
WO2013074668A1 (en) 2011-11-17 2013-05-23 Medtronic Minimed, Inc. Radition protecting composition and methods for making and using it
US9421287B2 (en) 2011-11-17 2016-08-23 Medtronic Minimed, Inc. Methods for making an aqueous radiation protecting formulation
US8999720B2 (en) 2011-11-17 2015-04-07 Medtronic Minimed, Inc. Aqueous radiation protecting formulations and methods for making and using them
US9103727B2 (en) 2012-01-03 2015-08-11 The Regents Of The University Of Michigan Composite Fabry-Pérot sensor
US10258736B2 (en) 2012-05-17 2019-04-16 Tandem Diabetes Care, Inc. Systems including vial adapter for fluid transfer
WO2013177573A2 (en) 2012-05-25 2013-11-28 Medtronic Minimed, Inc. Foldover sensors and methods for making and using them
US9493807B2 (en) 2012-05-25 2016-11-15 Medtronic Minimed, Inc. Foldover sensors and methods for making and using them
US11020028B2 (en) 2012-05-25 2021-06-01 Medtronic Minimed, Inc. Foldover sensors and methods for making and using them
US9586182B2 (en) * 2012-06-26 2017-03-07 Fujifilm Manufacturing Europe Bv Process for preparing membranes
US20150209733A1 (en) * 2012-06-26 2015-07-30 Fujifilm Manufacturing Europe Bv Process for Preparing Membranes
WO2014008297A1 (en) 2012-07-03 2014-01-09 Medtronic Minimed, Inc. Analyte sensors and production thereof
WO2014022641A1 (en) * 2012-08-01 2014-02-06 Massachusetts Institute Of Technology Actuation and control of stamp deformation in microcontact printing
US10772540B2 (en) 2012-12-06 2020-09-15 Medtronic Minimed, Inc. Microarray electrodes useful with analyte sensors and methods for making and using them
US10194840B2 (en) 2012-12-06 2019-02-05 Medtronic Minimed, Inc. Microarray electrodes useful with analyte sensors and methods for making and using them
WO2014089276A1 (en) 2012-12-06 2014-06-12 Medtronic Minimed, Inc. Microarray electrodes useful with analyte sensors and methods for making and using them
US11266332B2 (en) 2013-01-22 2022-03-08 Medtronic Minimed, Inc. Muting glucose sensor oxygen response and reducing electrode edge growth with pulsed current plating
US10426383B2 (en) 2013-01-22 2019-10-01 Medtronic Minimed, Inc. Muting glucose sensor oxygen response and reducing electrode edge growth with pulsed current plating
WO2014116293A1 (en) 2013-01-22 2014-07-31 Medtronic Minimed, Inc. Muting glucose sensor oxygen response and reducing electrode edge growth with pulsed current plating
US9324571B2 (en) * 2013-03-13 2016-04-26 Applied Materials, Inc. Post treatment for dielectric constant reduction with pore generation on low K dielectric films
US9962486B2 (en) 2013-03-14 2018-05-08 Tandem Diabetes Care, Inc. System and method for detecting occlusions in an infusion pump
WO2015069692A2 (en) 2013-11-07 2015-05-14 Medtronic Minimed, Inc. Enzyme matrices for use with ethylene oxide sterilization
US20170108464A1 (en) * 2013-12-10 2017-04-20 Life Technologies Corporation System and method for forming microwells
US9984891B2 (en) * 2016-04-21 2018-05-29 Shin-Etsu Chemical Co., Ltd. Method for forming organic film and method for manufacturing substrate for semiconductor apparatus
EP3263712A1 (en) 2016-06-29 2018-01-03 Roche Diabetes Care GmbH Galvanically functionalized sensors
US12613213B2 (en) 2016-06-29 2026-04-28 Roche Diabetes Care, Inc. Galvanically functionalized sensors
WO2018002108A1 (en) 2016-06-29 2018-01-04 Roche Diabetes Care Gmbh Galvanically functionalized sensors
WO2018170363A1 (en) 2017-03-17 2018-09-20 Medtronic Minimed, Inc. Metal pillar device structures and methods for making and using them in electrochemical and/or electrocatalytic applications
US12507915B2 (en) 2017-03-17 2025-12-30 Medtronic Minimed, Inc. Metal pillar device structures and methods for making and using them in electrochemical and/or electrocatalytic applications
EP3596452B1 (en) * 2017-03-17 2025-11-26 Medtronic MiniMed, Inc. Methods for making metal pillar device structures
EP4671752A2 (en) 2017-06-30 2025-12-31 Medtronic MiniMed, Inc. SENSOR INITIALIZATION METHODS FOR FASTER BODY SENSOR RESPONSE
WO2019005687A1 (en) 2017-06-30 2019-01-03 Medtronic Minimed, Inc. SENSOR INITIALIZATION METHODS FOR FASTER BODY SENSOR RESPONSE
US11069567B2 (en) * 2017-08-31 2021-07-20 International Business Machines Corporation Modulating metal interconnect surface topography
US10522398B2 (en) * 2017-08-31 2019-12-31 International Business Machines Corporation Modulating metal interconnect surface topography
US11382540B2 (en) 2017-10-24 2022-07-12 Dexcom, Inc. Pre-connected analyte sensors
US11350862B2 (en) 2017-10-24 2022-06-07 Dexcom, Inc. Pre-connected analyte sensors
US11331022B2 (en) 2017-10-24 2022-05-17 Dexcom, Inc. Pre-connected analyte sensors
US12150250B2 (en) 2017-10-24 2024-11-19 Dexcom, Inc. Pre-connected analyte sensors
US11706876B2 (en) 2017-10-24 2023-07-18 Dexcom, Inc. Pre-connected analyte sensors
US11943876B2 (en) 2017-10-24 2024-03-26 Dexcom, Inc. Pre-connected analyte sensors
WO2019147578A1 (en) 2018-01-23 2019-08-01 Medtronic Minimed, Inc. Implantable polymer surfaces exhibiting reduced in vivo inflammatory responses
WO2019156934A1 (en) 2018-02-07 2019-08-15 Medtronic Minimed, Inc. Multilayer electrochemical analyte sensors and methods for making and using them
WO2019157043A1 (en) 2018-02-08 2019-08-15 Medtronic Minimed, Inc. Glucose sensor electrode design
AU2019217883B2 (en) * 2018-02-08 2023-12-14 Medtronic Minimed, Inc. Methods for controlling physical vapor deposition metal film adhesion to substrates and surfaces
WO2019157106A2 (en) 2018-02-08 2019-08-15 Medtronic Minimed, Inc. Methods for controlling physical vapor deposition metal film adhesion to substrates and surfaces
CN111684098A (zh) * 2018-02-08 2020-09-18 美敦力泌力美公司 用于控制物理气相沉积金属膜与衬底和表面的粘附的方法
US12577648B2 (en) 2018-02-08 2026-03-17 Medtronic Minimed, Inc. Methods for controlling physical vapor deposition metal film adhesion to substrates and surfaces
WO2019157106A3 (en) * 2018-02-08 2019-10-10 Medtronic Minimed, Inc. Methods for controlling physical vapor deposition metal film adhesion to substrates and surfaces
US11220735B2 (en) 2018-02-08 2022-01-11 Medtronic Minimed, Inc. Methods for controlling physical vapor deposition metal film adhesion to substrates and surfaces
US11109784B2 (en) * 2018-04-09 2021-09-07 California Institute Of Technology Metal-enzyme sandwich layers
CN111742215A (zh) * 2018-04-09 2020-10-02 加州理工学院 金属-酶夹层
US11464908B2 (en) 2019-02-18 2022-10-11 Tandem Diabetes Care, Inc. Methods and apparatus for monitoring infusion sites for ambulatory infusion pumps
US11517229B2 (en) 2019-05-29 2022-12-06 Bionime Corporation Biosensing device and activation method thereof
EP3744248A1 (en) * 2019-05-29 2020-12-02 Bionime Corporation Biosensing device and activation method thereof
US11786154B2 (en) 2019-05-29 2023-10-17 Bionime Corporation Biosensing device and activation method thereof
WO2021021538A1 (en) 2019-07-26 2021-02-04 Medtronic Minimed, Inc. Methods to improve oxygen delivery to implantable sensors
WO2021021867A1 (en) 2019-08-01 2021-02-04 Medtronic Minimed, Inc. Micro-pillar working electrodes design to reduce backflow of hydrogen peroxide in glucose sensor
WO2022026542A1 (en) 2020-07-31 2022-02-03 Medtronic Minimed, Inc. Sensor identification and integrity check design
WO2022093574A1 (en) 2020-10-29 2022-05-05 Medtronic Minimed, Inc. Glucose biosensors comprising direct electron transfer enzymes and methods of making and using them
US12134068B2 (en) * 2021-01-04 2024-11-05 Saudi Arabian Oil Company High free volume membrane for gas separation
US20230277980A1 (en) * 2021-01-04 2023-09-07 Saudi Arabian Oil Company High free volume membrane for gas separation
EP4071251A1 (en) 2021-04-09 2022-10-12 Medtronic MiniMed, Inc. Hexamethyldisiloxane membranes for analyte sensors
US12517119B2 (en) 2021-04-19 2026-01-06 W. L. Gore & Associates, Inc. Electrochemical analyte biosensors, associated compositions of matter, and methods for making and using same
EP4134665A1 (en) 2021-08-13 2023-02-15 Medtronic MiniMed, Inc. Dry electrochemical impedance spectroscopy metrology for conductive chemical layers
EP4174188A1 (en) 2021-10-14 2023-05-03 Medtronic Minimed, Inc. Sensors for 3-hydroxybutyrate detection
EP4190908A1 (en) 2021-12-02 2023-06-07 Medtronic Minimed, Inc. Ketone limiting membrane and dual layer membrane approach for ketone sensing
CN114224334A (zh) * 2021-12-28 2022-03-25 苏州原理科技有限公司 植入式生物传感器及制备方法
EP4382611A1 (en) 2022-08-31 2024-06-12 Medtronic MiniMed, Inc. Sensors for 3-hydroxybutyrate detection
CN115736909A (zh) * 2022-12-08 2023-03-07 深圳清华大学研究院 血糖检测装置及血糖检测系统
US12376772B2 (en) * 2022-12-27 2025-08-05 Peking University Diabetes biosensor
US20250194970A1 (en) * 2022-12-27 2025-06-19 Peking University Diabetes biosensor

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US9541519B2 (en) 2017-01-10
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