WO2013144255A1 - Membrane d'espacement amélioré pour un capteur enzymatique in vivo - Google Patents

Membrane d'espacement amélioré pour un capteur enzymatique in vivo Download PDF

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Publication number
WO2013144255A1
WO2013144255A1 PCT/EP2013/056619 EP2013056619W WO2013144255A1 WO 2013144255 A1 WO2013144255 A1 WO 2013144255A1 EP 2013056619 W EP2013056619 W EP 2013056619W WO 2013144255 A1 WO2013144255 A1 WO 2013144255A1
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WO
WIPO (PCT)
Prior art keywords
hydrophilic
methacrylate
electrode system
copolymer
acrylate
Prior art date
Application number
PCT/EP2013/056619
Other languages
English (en)
Inventor
Arnulf Staib
Marcel Thiele
Karl-Heinz Koelker
Ewald Rieger
Original Assignee
F. Hoffmann-La Roche Ag
Roche Diagnostics Gmbh
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Priority claimed from PCT/EP2012/055406 external-priority patent/WO2012130841A1/fr
Priority to SI201330761T priority Critical patent/SI2831263T1/sl
Priority to DK13712795.7T priority patent/DK2831263T3/en
Priority to EP17169444.1A priority patent/EP3219807B1/fr
Priority to CA2867766A priority patent/CA2867766C/fr
Priority to RU2014142913A priority patent/RU2611038C2/ru
Priority to CN201380027747.XA priority patent/CN104334740B/zh
Priority to EP13712795.7A priority patent/EP2831263B1/fr
Priority to PL17169444T priority patent/PL3219807T3/pl
Priority to ES13712795.7T priority patent/ES2637811T3/es
Priority to PL13712795T priority patent/PL2831263T3/pl
Priority to JP2015502344A priority patent/JP6374860B2/ja
Priority to BR112014021373-9A priority patent/BR112014021373B1/pt
Application filed by F. Hoffmann-La Roche Ag, Roche Diagnostics Gmbh filed Critical F. Hoffmann-La Roche Ag
Publication of WO2013144255A1 publication Critical patent/WO2013144255A1/fr
Priority to IL234190A priority patent/IL234190B/en
Priority to ZA2014/06247A priority patent/ZA201406247B/en
Priority to US14/486,402 priority patent/US20150005605A1/en
Priority to HK15107281.6A priority patent/HK1206793A1/xx
Priority to US18/194,967 priority patent/US20230258595A1/en

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Classifications

    • C—CHEMISTRY; METALLURGY
    • C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
    • C12Q—MEASURING OR TESTING PROCESSES INVOLVING ENZYMES, NUCLEIC ACIDS OR MICROORGANISMS; COMPOSITIONS OR TEST PAPERS THEREFOR; PROCESSES OF PREPARING SUCH COMPOSITIONS; CONDITION-RESPONSIVE CONTROL IN MICROBIOLOGICAL OR ENZYMOLOGICAL PROCESSES
    • C12Q1/00—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions
    • C12Q1/001—Enzyme electrodes
    • C12Q1/002—Electrode membranes
    • 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, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue
    • A61B5/1486—Measuring characteristics of blood in vivo, e.g. gas concentration, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue using enzyme electrodes, e.g. with immobilised oxidase
    • 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, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue
    • A61B5/1486—Measuring characteristics of blood in vivo, e.g. gas concentration, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue using enzyme electrodes, e.g. with immobilised oxidase
    • A61B5/14865—Measuring characteristics of blood in vivo, e.g. gas concentration, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue using enzyme electrodes, e.g. with immobilised oxidase invasive, e.g. introduced into the body by a catheter or needle or using implanted sensors
    • 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
    • G01N27/3271—Amperometric enzyme electrodes for analytes in body fluids, e.g. glucose in blood
    • G01N27/3272—Test elements therefor, i.e. disposable laminated substrates with electrodes, reagent and channels
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B2562/00—Details of sensors; Constructional details of sensor housings or probes; Accessories for sensors
    • A61B2562/02—Details of sensors specially adapted for in-vivo measurements
    • 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, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue
    • A61B5/14532—Measuring characteristics of blood in vivo, e.g. gas concentration, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue for measuring glucose, e.g. by tissue impedance measurement
    • C—CHEMISTRY; METALLURGY
    • C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
    • C12Q—MEASURING OR TESTING PROCESSES INVOLVING ENZYMES, NUCLEIC ACIDS OR MICROORGANISMS; COMPOSITIONS OR TEST PAPERS THEREFOR; PROCESSES OF PREPARING SUCH COMPOSITIONS; CONDITION-RESPONSIVE CONTROL IN MICROBIOLOGICAL OR ENZYMOLOGICAL PROCESSES
    • C12Q1/00—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions
    • C12Q1/001—Enzyme electrodes
    • C12Q1/005—Enzyme electrodes involving specific analytes or enzymes
    • C12Q1/006—Enzyme electrodes involving specific analytes or enzymes for glucose

Definitions

  • the present invention relates to an electrode system for measuring the concentration of an analyte under in-vivo conditions, comprising an electrode with immobilized enzyme molecules and an improved diffusion barrier that controls diffusion of the analyte from body fluid surrounding the electrode system to the enzyme molecules.
  • the present invention relates to an electrode system for measuring the concentration of an analyte under in-vivo conditions, comprising an electrode with immobilized enzyme molecules, optionally a diffusion barrier that controls diffusion of the analyte from the exterior of the electrode system to the enzyme molecule and an improved spacer membrane which forms at least a portion of the outer layer of the electrode system.
  • Sensors with implantable or insertable electrode systems facilitate measurements of physiologically significant analytes such as, for example, lactate or glucose in a patient's body.
  • the working electrodes of systems of this type have electrically conductive enzyme layers in which enzyme molecules are bound which release charge carriers by catalytic conversion of analyte molecules. In the process, an electrical current is generated as measuring signal whose amplitude correlates to the analyte concentration.
  • Such electrode systems are e.g. known from WO 2007/147475 and WO 2010/028708, the contents of which are herein incorporated by reference.
  • the working electrodes of the electrode system are provided with a diffusion barrier that controls the diffusion of the analyte to be determined from the body fluid or tissue surrounding the electrode system to the enzyme molecules that are immobilized in the enzyme layer.
  • the diffusion barrier of the electrode system is a solid solution of at least two different polymers, preferably of acrylates.
  • the polymers may be copolymers, e.g. copolymers of methyl methacrylate and hydroxyethyl methacrylate or copolymers of butyl methacrylate and hydroxyethyl methacrylate.
  • WO 2007/147475 discloses a diffusion barrier made from a polymer having a zwitterionic structure.
  • An example of such a polymer is poly(2- methacryloyloxyethyl phosphorylcholine-co-n-butylmethacrylate).
  • the zwitterionic polymer may be mixed with another polymer, for example polyurethane.
  • WO 2006/058779 discloses an enzyme-based sensor with a combined diffusion and enzyme layer comprising at least one polymer material, and particles carry an enzyme, wherein the particles are dispersed in the polymer material.
  • the polymer may comprise hydrophilic as well as hydrophobic polymer chain sequences, for example, the polymer may be a high or low water uptake polyether-polyurethane copolymer.
  • block copolymers having at least one hydrophilic block and at least one hydrophobic block as a diffusion layer is not disclosed.
  • EP-A-2 163 190 describes an electrode system for the measurement of an analyte concentration in-vivo comprising a counterelectrode with an electric conductor, and a working electrode with an electric conductor on which an enzyme layer comprising immobilized enzyme molecules is localized.
  • a diffusion barrier controls the diffusion of the analyte from surrounding body fluids to the enzyme molecules.
  • the diffusion barrier may comprise hydrophilized polyurethanes obtainable by polycondensation of 4,4'- methylene-bis-(cyclohexylisocyanate) and diol mixtures which may be polyethyleneglycol and polypropyleneglycol.
  • the hydrophilic polyurethane layer may be covered with a spacer, e.g.
  • a copolymer of butyl methacrylate and 2-methacryloyloxyethyl-phosphoryl choline The use of block copolymers having at least one hydrophilic block and at least one hydrophobic block as a diffusion layer is not disclosed.
  • the use of a hydrophilic copolymer of (meth)acrylic monomers comprising more than 50 mol-% hydrophilic monomers is not disclosed either.
  • a diffusion barrier consisting of a single block copolymer having at least one hydrophilic block and at least one hydrophobic block.
  • the hydrophilic and hydrophobic blocks are covalently linked to each other.
  • the blocks are (meth)acrylate polymer blocks.
  • a subject-matter of the present invention is an electrode system for measuring the concentration of an analyte under in-vivo conditions, comprising an electrode with immobilized enzyme molecules and a diffusion barrier that controls diffusion of the analyte from the exterior of the electrode system to the enzyme molecules, characterized in that the diffusion barrier comprises a block copolymer having at least one hydrophilic block and at least one hydrophobic block.
  • the diffusion barrier comprises a single, i.e. only one block copolymer having at least one hydrophilic block and at least one hydrophobic block, i.e. further polymers or copolymers are absent. More preferably, the diffusion barrier consists of a single block copolymer having at least one hydrophilic block and at least one hydrophobic block.
  • the electrode system of the present invention is suitable for insertion or implantation into a body, e.g. a mammalian body such as a human body.
  • the electrode system is adapted for measuring a desired analyte in body fluid and/or body tissue, e.g. in the extracellular space (interstitium), in blood or lymph vessels or in the transcellular space.
  • the inserted or implanted electrode system is suitable for short-term application, e.g. 3-14 days, or for long-term application, e.g. 6-12 months.
  • a desired analyte may be determined by continuous or discontinuous measurements.
  • the electrode system of the invention is preferably part of an enzymatic, non-fluidic (ENF) sensor, wherein enzymatic conversion of the analyte is determined.
  • the sensor comprises a working electrode with immobilized enzyme molecules for the conversion of the analyte which results in the generation of an electrical signal.
  • the enzymes may be present in a layer covering the electrode. Additionally, redox mediators and/or electro-catalysts as well as conductive particles and pore formers may be present. This type of electrode is described e.g. in WO 2007/147475, the content of which is herein incorporated by reference.
  • the area of the working electrode is the sensitive area of the sensor.
  • This sensitive area is provided with a diffusion barrier that controls diffusion of the analyte from the exterior, e.g. body fluid and/or tissue surrounding the electrode system to the enzyme molecules.
  • the diffusion barrier can, for example, be a cover layer covering the enzyme layer, i.e. an enzyme-free layer.
  • diffusion-controlling particles are incorporated into the enzyme layer to serve as a diffusion barrier.
  • pores of the enzyme layer can be filled with the polymer which controls the diffusion of analyte molecules.
  • the thickness of the diffusion barrier is usually from about 2-20 pm, e.g. from about 2-15 ⁇ , or from about 5-20 ⁇ , particularly from about 5-10 pm or from about 10-15 pm (in dry state).
  • the diffusion barrier of the electrode system of the present invention comprises a block copolymer, preferably a single block copolymer having at least one hydrophilic block and at least one hydrophobic block.
  • the block copolymer may comprise an alternating sequence of blocks, i.e. a hydrophilic block is linked to a hydrophobic block.
  • the hydrophilic and hydrophobic blocks are covalently linked to each other within a polymer molecule.
  • the average molecular weight of the polymer (by weight) is usually from 20-70 kD, particularly from 25-60 kD and more particularly from 30-50 kD.
  • the molar ratio of the hydrophilic to hydrophobic portions in the block copolymer is usually in the range from about 75% (hydrophilic) : 25% (hydrophobic) to about 25% (hydrophilic) : 75% (hydrophobic), in the range from about 65% (hydrophilic) : 35% (hydrophobic) to about 35% (hydrophilic) : 65% (hydrophobic) or in the range from about 60% (hydrophilic) : 40% (hydrophobic) to about 40% (hydrophilic) : 60% (hydrophobic).
  • a hydrophilic block of the block copolymer consists of at least 90%, at least 95% and particularly completely of hydrophilic monomeric units. It usually has a length of from 50-400, e.g. from 50-200, or from 150-300 particularly from 100-150, or from 200-250 monomeric molecules.
  • a hydrophobic block of the copolymer consists of at least 90%, more particularly at least 95% and even more particularly completely of hydrophobic monomeric units. It has usually a length of from 50-300, e.g. from 50-200, or from 150-250, particularly from 80-150, or from 170-200 monomeric units.
  • hydrophilic blocks and/or the hydrophobic blocks preferably consist of (meth)acrylic-based units. More preferably, both the hydrophilic blocks and the hydrophobic blocks consist of (meth)acrylic-based monomeric units.
  • the hydrophilic monomeric units of the hydrophilic block are preferably selected from hydrophilic (meth)acryl esters, i.e. esters with a polar, i.e. OH, OCH 3 or OC2H5 group within the alcohol portion of the ester, hydrophilic (meth)acrylamides with an amide (NH 2 ) or an N-alkyl- or N,N-dialkylamide group, wherein the alkyl group comprises 1-3 C-atoms and optionally hydrophilic groups such as OH, OCH3 or OC 2 H 5 , and suitable (meth)acrylic units having a charged, e.g. an anionic or cationic group, such as acrylic acid (acrylate) or methacrylic acid (methacrylate). Further, combinations of monomeric units may be employed. Specific examples of preferred monomeric units for the hydrophilic block are selected from:
  • Iky I comprises
  • hydrophilic monomers are 2-hydroxyethyl methacrylate (HEMA) and/or 2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA). More preferably, the hydrophilic block consists of at least two different hydrophilic monomeric units. For example, it may be a random copolymer of at least two different hydrophilic monomeric units such as HEMA and 2-HPMA. ln order to introduce ionic groups into the monomer, charged monomeric units such as acrylic acid (acrylate) and/or methacrylic acid (methacrylate) may be incorporated into the hydrophilic block.
  • the hydrophilic block can be made from at least one non-ionic hydrophilic monomeric unit (e.g.
  • the hydrophilic block comprises an ionic monomeric unit such as acrylic acid or methacrylic acid, copolymerization with a hydrophilic monomer selected from the group of (meth)acrylamides, particularly ⁇ , ⁇ -dialkyl acryl- or methacrylamides is preferred.
  • the hydrophobic monomeric units of the hydrophobic block are preferably selected from hydrophobic acrylic and/or methacrylic units, styrene-based monomeric units or combinations thereof.
  • the hydrophobic monomeric units are selected from hydrophobic (meth)acryl esters, e.g. esters having an alcohol portion with 1 -3 C-atoms without hydrophilic group.
  • Specific examples of monomeric units for the hydrophobic block are selected from: methyl acrylate,
  • MMA methyl methacrylate
  • EMA ethyl methacrylate
  • BUMA n-butyl methacrylate
  • the hydrophobic block preferably comprises at least two different hydrophobic monomeric units, which are e.g. present as a random copolymer.
  • the hydrophobic block comprises methyl methacrylate (MMA) and n-butyl methacrylate (BUMA).
  • the hydrophobic block is a random copolymer of MMA and BUMA.
  • the molar ratio between MMA and BUMA is preferably about 60% (MMA) : 40% (BUMA) to about 40% (MMA) : 60% (BUMA), e.g. about 50% (MMA) : 50% (BUMA).
  • the glass transition temperature of the hydrophobic block is preferably 100°C or less, 90°C or less or 80°C or less, e.g. about 40-80°C.
  • the hydrophobic block may consist of styrenic units, e.g. of polystyrene having a glass transition temperature of about 95°C.
  • the block copolymers used in the present invention may be manufactured according to known methods (Boker et al., Macromolecules 34 (2001 ), 7477- 7488).
  • the block copolymers may be applied to the electrode system by usual techniques, e.g. by providing a solution of the block copolymer in a suitable solvent or solvent mixture, e.g. an organic solvent, such as ether, which is applied to the prefabricated electrode system and dried thereon.
  • a suitable solvent or solvent mixture e.g. an organic solvent, such as ether
  • the block copolymer When the block copolymer is contacted with water, it shows a water uptake of preferably about 15%-30% by weight (based on the polymer dry weight) at a temperature of 37°C and a pH of 7.4 (aqueous phosphate buffer 10 mM KH2PO4, 10 mM NaH 2 P0 4 and 147 mM NaCI).
  • the diffusion barrier may also comprise further components, particularly non-polymeric components, which may be dispersed and/or dissolved in the polymer.
  • these further compounds include plasticizers, particularly biocompatible plasticizers, such as tri-(2-ethylhexyl) trimellitate and/or glycerol.
  • the diffusion barrier of the invention has a high effective diffusion coefficient D ef f for glucose which is preferably > 10 10 cm 2 /s, more preferably > 5 10 0 cm 2 /s, and even more preferably > 10 "9 cm 2 /s, and e.g. up to 10 7 or 10 8 cm 2 /s at a temperature of 37°C and a pH of 7.4.
  • the effective diffusion coefficient is preferably determined as described in Example 4 according to the equation:
  • SE m is the sensitivity of the working electrode
  • F is the area of the working electrode
  • SE and L m may be determined as described in the Examples.
  • the electrode system of the present invention is suitable for measuring the concentration of an analyte under in- vivo conditions, i.e. when inserted or implanted into a body.
  • the analyte may be any molecule or ion present in tissue or body fluid, for example oxygen, carbon dioxide, salts (cations and/or anions), fats or fat components, carbohydrates or carbohydrate components, proteins or protein components, or other type of biomolecules.
  • body fluid e.g. blood and tissue such as oxygen, carbon dioxide, sodium cations, chloride anions, glucose, urea, glycerol, lactate and pyruvate.
  • the electrode system comprises an enzyme immobilized on an electrode.
  • the enzyme is suitable for the determination of a desired analyte.
  • the enzyme is capable of catalytically converting the analyte and thereby generating an electric signal detectable by the electric conductor of the working electrode.
  • the enzyme for measuring the analyte is preferably an oxidase, for example glucose oxidase or lactate oxidase or a dehydrogenase, for example a glucose dehydrogenase or a lactate dehydrogenase.
  • the enzyme layer may also comprise an electrocatalyst or a redox mediator which favours the transfer of electrons to conductive components of the working electrode, e. g.
  • Suitable electro-catalysts are metal oxides such as manganese dioxide or organo-metallic compounds such as cobalt phthalo-cyanine.
  • the redox mediator is capable of degrading hydrogen peroxide thereby counteracting depletion of oxygen in the surroundings of the working electrode.
  • a redox mediator may be covalently bound to the enzyme and thereby effect direct electron transfer to the working electrode.
  • Suitable redox mediators for direct electron transfer are prosthetic groups, such as pyrrolo quinoline quinone (PQQ), flavine adenine dinucleotide (FAD) or other known prosthetic groups. Enzymes immobilized on electrodes are e.g. described in WO 2007/147475, the content of which is herein incorporated by reference.
  • a preferred embodiment of the electrode system comprises a counterelectrode with an electrical conductor and a working electrode with an electrical conductor on which an enzyme layer and the diffusion barrier are arranged.
  • the enzyme layer is preferably designed in the form of multiple fields that are arranged on the conductor of the working electrode at a distance, e.g. at least 0.3 mm or at least 0.5 mm from each other.
  • the individual fields of the working electrode may form a series of individual working electrodes. Between these fields, the conductor of the working electrode may be covered by an insulation layer.
  • the electrode system of the invention may additionally comprise a reference electrode capable of supplying a reference potential for the working electrode, e.g. an Ag/Ag-CI reference electrode.
  • a reference electrode capable of supplying a reference potential for the working electrode, e.g. an Ag/Ag-CI reference electrode.
  • an electrode system according to the invention can have additional counter- and/or working electrodes.
  • the electrode system may be part of a sensor, e.g. by being connected to a potentiostat and an amplifier for amplification of measuring signals of the electrode system.
  • the sensor is preferably an enzymatic non-fluidic (ENF) sensor, more preferably an electrochemical ENF sensor
  • the electrodes of the electrode system may be arranged on a substrate that carries the potentiostat or be attached to a circuit board that carries the potentiostat.
  • a further subject-matter of the invention is related to the use of a block copolymer having at least one hydrophilic block and at least one hydrophobic block as a diffusion barrier for an enzymatic electrode.
  • the block copolymer is preferably as described above, e.g. a single block-copolymer.
  • the diffusion barrier and the enzymatic electrode are preferably also as described above.
  • Fig. 1 shows an exemplary embodiment of an electrode system according to the invention.
  • Fig. 2 shows a detail view of Fig. 1.
  • Fig. 3 shows another detail view of Fig. 1.
  • Fig. 4 shows a section along the section line CC of Fig. 2.
  • Fig. 5 shows the sensitivity (with standard deviation) of four glucose sensors (at 10 mM glucose) provided with different block polymers (C, F, D, B) as barrier layers.
  • Fig. 6 shows the sensor drift of four glucose sensors provided with different block copolymers (A, C, D, F) as barrier layers.
  • Fig. 7 shows the conductivity of block copolymer A dependent on time (2 experiments).
  • Fig. 8 shows the conductivity of block copolymer F dependent on time
  • Fig. 9 shows the conductivity of block copolymer H dependent on time for a layer thickness of 2.77 pm or 4.43 pm, respectively.
  • Fig. 10 shows the fibrinogen adhesion to different spacer membrane polymers in-vitro (Adapt® and Eudragit E100) with respect to an uncoated plate (Blank).
  • Fig. 13 shows the secretion of cytokine IL-8 by THP-1 cells after incubation with tissue culture plates coated with a spacer membrane (Adapt®, Lipidure CM5206 and Eudragit E100) or uncoated (Polyst.), and an additional layer of fibrinogen.
  • Figure 1 shows an exemplary embodiment of an electrode system for insertion into body tissue of a human or animal, for example into cutis or subcutaneous fatty tissue.
  • a magnification of detail view A is shown in Figure 2
  • a magnification of detail view B is shown in Figure 3.
  • Figure 4 shows a corresponding sectional view along the section line, CC, of Figure 2.
  • the electrode system shown has a working electrode 1 , a counterelectrode 2, and a reference electrode 3. Electrical conductors of the electrodes 1 a, 2a, 3a are arranged in the form of metallic conductor paths, preferably made of palladium or gold, on a substrate 4.
  • the substrate 4 is a flexible plastic plate, for example made of polyester.
  • the substrate 4 is less than 0.5 mm thick, for example 100 to 300 micrometers, and is therefore easy to bend such that it can adapt to movements of surrounding body tissue after its insertion.
  • the substrate 4 has a narrow shaft for insertion into body tissue of a patient and a wide head for connection to an electronic system that is arranged outside the body.
  • the shaft of the substrate 4 preferably is at least 1 cm in length, in particular 2 cm to 5 cm.
  • one part of the measuring facility projects from the body of a patient during use.
  • the working electrode 1 carries an enzyme layer 5 that contains immobilized enzyme molecules for catalytic conversion of the analyte.
  • the enzyme layer 5 can be applied, for example, in the form of a curing paste of carbon particles, a polymeric binding agent, a redox mediator or an electro-catalyst, and enzyme molecules. Details of the production of an enzyme layer 5 of this type are disclosed, for example, in WO 2007/147475, reference to which is be made in this context.
  • the enzyme layer 5 is not applied continuously on the conductor 1 a of the working electrode 1 , but rather in the form of individual fields that are arranged at a distance from each other.
  • the individual fields of the enzyme layer 5 in the exemplary embodiment shown are arranged in a series.
  • the conductor 1 a of the working electrode 1 has narrow sites between the enzyme layer fields that are seen particularly well in Figure 2.
  • the conductor 2a of the counterelectrode 2 has a contour that follows the course of the conductor 1 a of the working electrode 1. This means results in an intercalating or interdigitated arrangement of working electrode 1 and counterelectrode 2 with advantageously short current paths and low current density.
  • the counterelectrode 2 can be provided with a porous electrically conductive layer 6 that is situated in the form of individual fields on the conductor 2a of the counterelectrode 2.
  • this layer 6 can be applied in the form of a curing paste of carbon particles and a polymeric binding agent.
  • the fields of the layer 6 preferably have the same dimensions as the fields of the enzyme layer 5, although this is not obligatory.
  • measures for increasing the surface of the counterelectrode can just as well be foregone and the counterelectrode 2 can just as well be designed to be a linear conductor path with no coatings of any kind, or with a coating made from the described block copolymer and optionally a spacer.
  • the reference electrode 3 is arranged between the conductor 1 a of the working electrode 1 and the conductor 2a of the counterelectrode 2.
  • the reference electrode shown in Figure 3 consists of a conductor 3a on which a field 3b of conductive silver/silver chloride paste is arranged.
  • Figure 4 shows a schematic sectional view along the section line, CC, of Figure 2.
  • the section line, CC extends through one of the enzyme layer fields 5 of the working electrode 1 and between the fields of the conductive layer 6 of the counterelectrode 2.
  • the conductor 1 a of the working electrode 1 can be covered with an electrically insulating layer 7, like the conductor 2a of the counterelectrode 2 between the fields of the conductive layers 6, in order to prevent interfering reactions which may otherwise be catalysed by the metal of the conductor paths 1 a, 2a.
  • the fields of the enzyme layer 5 are therefore situated in openings of the insulation layer 7.
  • the fields of the conductive layer 6 of the counterelectrode 2 may also be placed on top of openings of the insulation layer 7.
  • the enzyme layer 5 is covered by a cover layer 8 which presents a diffusion resistance to the analyte to be measured and therefore acts as a diffusion barrier.
  • the diffusion barrier 8 consists of a single copolymer with alternating hydrophilic and hydrophobic blocks as described above.
  • a favourable thickness of the cover layer 8 is, for example, 3 to 30 ⁇ , particularly from about 5-10 ⁇ or from about 10-15 pm. Because of its diffusion resistance, the cover layer 8 causes fewer analyte molecules to reach the enzyme layer 5 per unit of time. Accordingly, the cover layer 8 reduces the rate at which analyte molecules are converted, and therefore counteracts a depletion of the analyte concentration in surroundings of the working electrode.
  • the cover layer 8 extends continuously essentially over the entire area of the conductor 1 a of the working electrode 1 .
  • a biocompatible membrane may be arranged as spacer 9 that establishes a minimal distance between the enzyme layer 5 and cells of surrounding body tissue. This means advantageously generates a reservoir for analyte molecules from which analyte molecules can get to the corresponding enzyme layer field 5 in case of a transient disturbance of the fluid exchange in the surroundings of an enzyme layer field 5. If the exchange of body fluid in the surroundings of the electrode system is transiently limited or even prevented, the analyte molecules stored in the spacer 9 keep diffusing to the enzyme layer 5 of the working electrode 1 where they are converted.
  • the spacer 9 therefore causes a notable depletion of the analyte concentration and corresponding falsification of the measuring results to occur only after a significantly longer period of time.
  • the membrane forming the spacer 9 also covers the counterelectrode 2 and the reference electrode 3.
  • the spacer membrane 9 can, for example, be a dialysis membrane.
  • a dialysis membrane is understood to be a membrane that is impermeable for molecules larger than a maximal size.
  • the dialysis membrane can be prefabricated in a separate manufacturing process and may then be applied during the fabrication of the electrode system.
  • the maximal size of the molecules for which the dialysis membrane is permeable is selected such that analyte molecules can pass, while larger molecules are retained.
  • a coating made of a polymer that is highly permeable for the analyte and water, for example on the basis of polyurethane or of acrylate, can be applied over the electrode system as spacer membrane 9.
  • the spacer is made from a copolymer of (meth)acrylates.
  • the spacer membrane is a copolymer from at least 2 or 3 (meth)acrylates. More preferably, the spacer membrane comprises more than 50 mol-%, at least 60 mol-% or at least 70 mol-% hydrophilic monomer units, e.g. HEMA and/or 2-HPMA, and up to 40 mol-% or up to 30 mol-% hydrophobic units, e.g. BUMA and/or MMA.
  • the spacer may be a random or block copolymer.
  • An especially preferred spacer membrane comprises MMA or BUMA as hydrophobic moieties and 2-HEMA and/or 2-HPMA as hydrophilic moieties.
  • the amount of the hydrophilic monomers HEMA and/or HPMA is between 80 mol-% to 85 mol-% and the amount of hydrophobic component MMA and/or BUMA is between 15 mol-% and 20 mol-%.
  • the very preferred spacer membrane of the invention is made of the copolymer Adapt® (Biolnteractions Ltd, Reading, England).
  • Adapt® comprises BUMA as hydrophobic moiety and 2-HEMA and 2-HPMA as hydrophilic moieties, wherein the amount of the 2-HEMA hydrophilic monomers is about 80 mol-%.
  • the spacer membrane is highly permeable for the analyte, i.e. it does significantly lower the sensitivity per area of the working electrode, for example 20% or less, or 5% or less with a layer thickness of less than about 20 ⁇ , preferably less than about 5 ⁇ .
  • An especially preferred thickness of the spacer membrane is from about 1 to about 3 pm.
  • the enzyme layer 5 of the electrode system can contain metal oxide particles, preferably manganese dioxide particles, as catalytic redox mediator.
  • Manganese dioxide catalytically converts hydrogen peroxide that is formed, for example, by enzymatic oxidation of glucose and other bioanalytes.
  • the manganese di-oxide particles transfer electrons to conductive components of the working electrode 1 , for example to graphite particles in the enzyme layer 5.
  • the catalytic degradation of hydrogen peroxide counteracts any decrease of the oxygen concentration in the enzyme layer 5.
  • this allows the conversion of the analyte to be detected in the enzyme layer 5 to not be limited by the local oxygen concentration.
  • the use of the catalytic redox mediator therefore counteracts a falsification of the measuring result by the oxygen concentration being low.
  • Another advantage of a catalytic redox mediator is that it prevents the generation of cell-damaging concentrations of hydrogen peroxide.
  • the preferred spacer membrane polymer described herein may be used as an outer coating for a diffusion barrier of the present invention, but also as an outer coating of an electrode system in general, particularly of an electrode system for measuring the concentration of an analyte under in-vivo conditions, comprising an electrode with immobilized enzyme molecules and a diffusion barrier that controls diffusion of the analyte from the exterior of the electrode system to the enzyme molecules.
  • the spacer membrane can be arranged on the diffusion barrier, however, the spacer membrane can also be arranged directly on the enzyme layer. In this last context, the spacer membrane can also act as a diffusion barrier itself and slow down the diffusion of analyte molecules to the enzyme layer.
  • the spacer membrane When the electrode system of the invention is inserted or implanted into a body, the spacer membrane is the interface between the implanted sensor and the surrounding body fluid or tissue. Consequently, when exposed to the body fluid or tissue, the spacer membrane of the invention must be mechanically robust so that it is neither deformed nor shoved off the sensor. To this end, the spacer membrane copolymer water uptake and the concomitant swelling of the copolymer must be limited, albeit the inherent hydrophilicity of the copolymer.
  • the relative water uptake of the spacer membrane copolymer should not exceed 50 wt-%, preferably 40 wt-%, more preferably 30 wt-% based on the total rate of the copolymer.
  • the measurement of the relative water uptake is performed by subjecting the dry copolymer to an excess of phosphate-buffer (pH 7.4) for 48 h at a temperature of 37 °C.
  • the relative water uptake (WU%) is preferably determined according to the equation:
  • WU% (m 2 -m 1 )/m 1 x 100, wherein mi and m 2 represent, respectively, the mass of the dry copolymer and the copolymer after hydration according to the above measurement conditions.
  • the inventors of the present invention determined that the preferred spacer membrane made of the copolymer Adapt® takes up 33 ⁇ 1 .8 wt-% of phosphate-buffer (pH 7.4) relative to its own weight over 48 h at 37° C. Under the same conditions, a membrane of polymer Lipidure CM5206 (NOF Corporation, Japan) takes up 157 ⁇ 9.7 wt-% of phosphate-buffer relative to its own weight.
  • the lower water uptake of the polymer advantageously increases the mechanical stability of the spacer membrane of the present invention.
  • Lipidure CM5206 shows a higher water uptake and swells to a hydrogel which is more fragile, easily deformable or shoved off, particularly when applied on an electrode system of an enzymatic in-vivo sensor.
  • the spacer is in direct contact with the tissue and/or the body fluid, like interstitial fluid or blood, containing biomolecules like proteins and cells.
  • the spacer membrane must protect the inserted and implanted sensor in the tissue and/or body fluid environment and, thus, minimize the tissue reaction of the body to the implant.
  • reactions of the body against implanted material are known as "foreign body response" (FBR).
  • FBR the body tries to destroy the implant or, if it is not possible, to create a capsule to separate it from the surrounding tissue (foreign body granuloma).
  • the first step of the FBR reaction is binding of proteins (e.g.
  • fibrinogen contains a structural motif that binds to the monocyte receptor MAC-1 .
  • fibrinogen binds to the surface of the implant, it changes its conformation and exposes the binding site for MAC-1. Consequently, immune cells, like monocytes, are recruited to the implant and activated, secreting enzymes and radicals to attack the implant. Additionally, immune cells secrete soluble factors, i.e. cytokines, to recruit and activate other immune cells and thereby amplifying the immune response.
  • a fibrous capsule is formed by connective tissue cells and proteins. This capsule, however, is a diffusion barrier for analytes to reach the sensor.
  • an improved spacer membrane on an electrode-system of an enzymatic in-vivo sensor further provides the reduction of the tissue response to the implant and inhibits the formation of a capsule separating the sensor from the surrounding tissue and body fluids.
  • an electrode system for measuring the concentration of an analyte under in-vivo conditions comprising an electrode with immobilized enzyme molecules and preferably a diffusion barrier that controls diffusion of the analyte form the exterior of the electrode system to the enzyme molecules, characterised in that a spacer membrane forms at least a portion of the outer layer of the electrode system, wherein the spacer membrane comprises a hydrophilic copolymer of acrylic and/or methacrylic monomers, wherein the polymer comprises more than 50 mol-% hydrophilic monomers.
  • the spacer membrane of the invention does have limited protein-binding capacity to protect the electrode system of the sensor from protein adsorption that might trigger response of immune cells and might limit or interfere its performance in-vivo.
  • Example 5 and 6 show that the preferred spacer membrane of the invention provides little binding to fibrinogen and prevents conformational change of fibrinogen that would lead to the exposure of the MAC-1 binding motif for monocytes.
  • the spacer membrane copolymer material does not activate immune cells itself.
  • Example 6 it could be demonstrated that the spacer membrane copolymer of the invention is able to attenuate the activation of immune cells by the implanted sensor.
  • the spacer membrane is a biocompatible material, in particular, is compatible with body fluids, e.g. with blood.
  • Example 7 shows that the spacer membrane copolymer of the present invention is able to prevent haemolysis and the complement activation by the implanted sensor.
  • the spacer membrane of the invention advantageously not only shows a high mechanical stability, but also has optimal biocompatible properties, which is surprising due to the low water uptake when wetted.
  • the diffusion barrier is preferably as described herein, it may however also have a different composition or may be absent.
  • the diffusion barrier preferably comprises a block copolymer having at least one hydrophilic block and at least one hydrophobic block as described herein.
  • the diffusion barrier comprises hydrophilic polyurethanes.
  • the hydrophilic polyurethanes used as diffusion membrane can be produced by polyaddition of an (poly)diisocyanat, preferably 4-4-methylene-bis(cyclohexylisocyanate) with a polyalcohol, preferably a diol mixture.
  • the components of the diol mixture are preferably polyalkylene glycols, such as polyethylene glycol (PEG) and polypropylene glycol (PPG) and aliphatic diols, such as ethylene glycol.
  • the hdydrophilic polyurethane comprises 45-55 mol-%, preferably 50 mol-% isocyanate and 25-35 mol-%, preferably 30 mol-% ethylene glycol.
  • the degree of hydrophilization is then adjusted by the ratio of PEG to PPG.
  • the polyurethane comprises 2-3 mol-%, more preferably 2.5 mol-% PEG and 17-18 mol-%, preferably 17,5 mol-% PPG.
  • the proportion of PEG may be increased, for instance, to 4.5 - 5.5 mol-%, preferably 5 mol-% PEG, in order to obtain an extremely hydrophilic polyurethane.
  • the different hydrophilic variants of the polyurethanes may also be mixed in order to optimize the properties of the diffusion barrier.
  • the preferred acrylic and methacrylic monomers of the spacer membrane copolymer are as described herein.
  • the hydrophilic monomeric units are preferably selected from hydrophilic (meth)acryl esters, i.e. esters with a polar, i.e. OH, OCH3 or OC 2 H 5 group within the alcohol portion of the ester, hydrophilic (meth)acrylamides with an amide (NH 2 ) or an N-alkyl- or ⁇ , ⁇ -dialkylamide group, wherein the a Iky I group comprises 1 -3 C-atoms and optionally hydrophilic groups such as OH, OCH 3 or OC 2 H5, and suitable (meth)acrylic units having a charged, e.g. an anionic or cationic group, such as acrylic acid (acrylate) or methacrylic acid (methacrylate). Further, combinations of monomeric units may be employed.
  • hydrophilic (meth)acryl esters i.e. esters with a polar, i.e. OH, OCH3 or OC 2 H 5 group within the alcohol portion of the ester
  • preferred monomeric units for the hydrophilic block are selected from:
  • alkyl comprises
  • Preferred hydrophilic monomers are 2-hydroxyethyl methacrylate (HEMA) and/or 2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA).
  • HEMA 2-hydroxyethyl methacrylate
  • 2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA).
  • the hydrophobic monomeric units are preferably selected from hydrophobic acrylic and/or methacrylic units or combinations thereof.
  • the hydrophobic monomeric units are selected from hydrophobic (meth)acryl esters, e.g. esters having an alcohol portion with 1-3 C-atoms without hydrophilic group.
  • monomeric units for the hydrophobic block are selected from: methyl acrylate,
  • MMA methyl methacrylate
  • EMA ethyl methacrylate
  • BUMA n-butyl methacrylate
  • the hydrophobic block comprises methyl methacrylate (MMA) and n-butyl methacrylate (BUMA).
  • the outer spacer membrane preferably covers at least the working electrode portion comprising the enzyme molecules and optionally also other portions, e.g. the counter electrode. If one is present, the spacer membrane also covers the reference electrode.
  • the spacer membrane preferably covers the entire implanted surface of the electrode system.
  • the spacer membrane preferably covers the working electrode, optionally the counter-electrode and the reference electrode if present in the form of a continuous layer.
  • the electrode system comprising the improved spacer membrane of the invention may be part of a sensor, e.g. by being connected to a potentiostat and an amplifier for amplification of measuring signals of the electrode system.
  • the sensor is preferably an enzymatic non-fluidic (ENF) sensor, more preferably an electrochemical ENF sensor
  • the electrodes of the electrode system may be arranged on a substrate that carries the
  • the potentiostat or be attached to a circuit board that carries the potentiostat.
  • the sensor is for the measurement of glucose.
  • a further subject-matter of the invention is related to the use of a hydrophilic copolymer of acrylic and/or methacrylic monomers, wherein the hydrophilic compolymer comprises more than 50 mol-% hydrophilic monomers as a spacer membrane for an enzymatic electrode.
  • the hydrophilic copolymer is preferably as described above.
  • the spacer membrane is used for minimising the foreign body reaction (FRB) against the enzymatic electrode when it is inserted or implanted into the body.
  • FAB foreign body reaction
  • Example 1 Permeability of an enzymatic non-fluidic (ENF) glucose sensor with distributed electrodes for transcutaneous implantation having a diffusion layer consisting of one single block copolymer.
  • the sensor was built on a prefabricated palladium strip conductor structure on a polyester substrate having a thickness of 250 ⁇ .
  • Working electrode (WE) and counterelectrode (CE) were arranged distributedly (as shown in Figs 1-2).
  • the fields of the CE were overprinted with carbon paste, the rest of the strip conductor was insulated.
  • the fields of the WE were overprinted with a mixture of cross-linked glucose oxidase (enzyme), conductive polymer paste and electric catalyst, here manganese(IV)-oxide (Technipur).
  • the remaining paths of the strip conductor were again insulated.
  • the reference electrode (RE) consists of Ag/AgCI paste. The electrodes cover about 1 cm of the sensor shaft.
  • the WE-fields were coated with a block copolymer diffusion layer consisting of a HEMA block and a BUMA block.
  • the thickness of the layer is 7 pm.
  • Fig. 5 shows the sensor sensitivity with standard deviations for the four different diffusion layers.
  • the sensors having a diffusion layer of block copolymer B are an exception. Even though polymer B has a relative ratio of hydrophobic to hydrophilic amount similar to polymer F, the sensitivity and thus the permeability for glucose is reduced. Empirically it can be stated that in case of polymer B the total chain length - corresponding to the molecular weight (total) of the copolymer molecule - is so large that the permeability of the layer is reduced. This may also be seen in the gravimetrically determined water uptake of block copolymer B as compared to the remaining polymers. Polymer B has a water uptake of 10.6% ⁇ 1 .5% (weight percent referred to the polymer dry weight). Polymer C lies at 15.6% ⁇ 0.0%, polymer F at 16.5 ⁇ 3.1 % and polymer D at 27% ⁇ 1 .7%. Example 2 Mechanic flexibility of the diffusion layer of an ENF glucose sensor
  • the sensor was manufactured as described in WO2010/028708, however having a diffusion layer according to the present invention. It was assumed that the glass transition temperature (Tg) is a substitute parameter for the mechanic flexibility. Further, it was assumed that the glass transition temperature, which may be allocated to the hydrophobic block, determines the mechanic flexibility in in-vivo applications. It should be noted that several Tgs may be identified for one block copolymer, corresponding to the number of blocks.
  • the sensors were coated with the same electrode pastes as in Example 1. Then, some of the sensors were coated with a copolymer selected from MMA-HEMA (produced by Polymer Source, Montreal). This polymer (called E) has a total molecular weight of 41 kD, the molar ratio of MMA (hydrophobic amount) to HEMA is 60%:40%. The glass transition temperature of the hydrophobic block is 1 1 1 °C, determined by DSC and a heating rate of 10°C/min.
  • MMA-HEMA produced by Polymer Source, Montreal
  • A a diffusion layer of a block copolymer of the invention
  • the hydrophobic block of said copolymer A contains MMA and BUMA at equal molar amounts in a randomised sequence. Again, the molar ratio of the hydrophobic part to the hydrophilic part is 60%:40%. The molecular weight is 36 kD.
  • the Tg of the hydrophobic block decreases, due to the randomized sequence of MMA and BUMA (Tg about 45°C), to 73°C.
  • Both diffusion layers were generated from the respective solution (25%) of the copolymers in ether and dried as in Example 1 .
  • the thickness of the diffusion layers was 7 pm.
  • a spacer layer was applied subsequently via dip coating and dried 24h at room temperature.
  • the spacer layer was made of Lipidure CM 5206, produced by NOF Japan.
  • sensors having a copolymer E diffusion layer show sporadic cracks in the diffusion layer. This is taken as an effect of the mechanic load.
  • sensors having a copolymer A diffusion layer do not show any cracks under identical load. This is obviously due to the reduction of Tg, which increases the mechanic stability of the copolymer.
  • a physical mixture of two copolymers, as disclosed in WO2010/028708, is no longer required.
  • Example 3 Optimized permeation behaviour of an ENF glucose sensor with distributed electrode and diffusion layer according to the invention.
  • the sensors were provided with a spacer layer as described in Example 2.
  • the sensor was connected with a measuring system on the sensor head, which transfers the measured data to a data store.
  • the in-vitro measurements were carried out as in Example 1 , however over a measuring period of 7 days. From the measured data, the sensitivity drift was calculated over the respective measuring period for each sensor.
  • Figure 6 shows for each sensor variant, i.e. sensors of a variant of the diffusion layer, the mean value of the in- vitro drift value for the group.
  • the initial phase of the measurement - the first 6h, the so-called startup phase - was excluded from the calculation.
  • copolymer A With the hydrophobic block of a random copolymer of MMA and BUMA, leads to a very low, slightly negative, drift.
  • Fig. 7 shows that the conductivity of copolymer A remained nearly constant after a short startup phase.
  • a sensor having block copolymer A shows a negligible drift, which is due to a very low permeability alteration in the conductivity measurement.
  • a strong increase of conductivity is observed in copolymer A.
  • a very fast startup is observed, which is terminated after about 1 hour.
  • the diffusion layer is completely wetted and has terminated its structural reorganisation due to water uptake.
  • the extent of the structural change presumably depends on the Tg. It seems plausible that a copolymer having an increased Tg passes a reorganisation, which is limited in time and amplitude, as compared to a copolymer having a Tg in the range of the ambient temperature.
  • sensors with copolymer A show a comparatively high sensitivity at the start of measurements as compared to sensors having a copolymer F diffusion layer. This is to be expected due to the identical relative ratios between hydrophobic and hydrophilic blocks.
  • the achieved sensitivity range of 1 to 1.5 nA/mM (see Example 1 ) is deemed ideal. This sensitivity is likewise obtained for sensors having a diffusion layer consisting of copolymer A.
  • an optimal sensor may preferably be obtained with a diffusion layer of a block copolymer, having a hydrophobic block with at least two different randomly arranged hydrophobic monomeric units, such as block copolymer A. None of the other block copolymers, whose hydrophobic blocks only consist of a single monomeric unit reaches a quality, which could be compared in all three parameters with copolymer A.
  • a multiple field sensor (10 fields of working electrodes and counterelectrodes, respectively) for the continuous measurement of the glucose was produced and characterized in-vitro.
  • the sensor was provided with a diffusion layer consisting of a block copolymer comprising a hydrophobic block of random copolymerized methyl methacrylate (MMA) and n-butyl methacrylate (BUMA) and a hydrophilic block of 2-hydroxyethyl methacrylate (HEMA).
  • MMA methyl methacrylate
  • BUMA n-butyl methacrylate
  • HEMA 2-hydroxyethyl methacrylate
  • the molecular weights Mn of each block are separately indicated in the above Table 2 and represent average values, as polymers are known to have distributions of molecular chain lengths around a specified mean value. This also applies to the derived quantities in Table 2.
  • the indicated glass transition temperatures of the hydrophobic block are within the desired range in order to guarantee mechanical flexibility.
  • the decisive parameter with regard to the permeability of the diffusion barrier for the analyte is the sensitivity per area unit of the working electrode (i.e. the geometric area).
  • the sensitivity SE was calculated from current (I) measurements at 10 mM and at 0 mM glucose concentration in phosphate- buffered solution (pH 7.4) in nA/mM:
  • the linearity Y of the in-vitro function curve is an indication of the diffusion control functionality of the polymer cover layer on the working electrode. It was calculated from current measurements at 20 mM, 10 mM and 0 mM glucose concentration in %:
  • Y 20mM 50 [l(20mM) - l(0mM)]/[l(10mM) - l(OmM)] for each of the analyzed sensors. From the individual measurement values the mean linearity value and its standard deviation were determined (cf. Table 3).
  • the layer thickness L of the diffusion barrier of the sensors was determined by optical measurement for each of the polymers.
  • the corresponding mean values were computed for a sample of > 23 sensors with the same polymer.
  • the effective diffusion coefficient D e ff of the cover layer may be calculated: in cm 2 /s, wherein SE m and L m are the respective mean values for the sensitivity and the layer thickness, and F is the total area of all working electrode spots.
  • the diffusion coefficient was also determined with an alternative method, e.g. permeation of glucose from a chamber with a glucose solution into a chamber with a glucose-free buffer through a film of the polymer. According to this method, a similar value for the diffusion coefficient was obtained (1.17-10- 9 crr Vs).
  • Eudragit E100 is a cationic copolymer based on dimethylaminoethyl methacrylate, butyl methacrylate and methyl methacrylate. The polymers were dried over night at 40°C. Thereafter the spacer materials were overlaid with fibrinogen solution. The solution contained fibrinogen from human plasma that was conjugated to the fluorescent dye Alexa488 (purchased from Invitrogen).
  • the fibrinogen solution was aspirated and the spacer layers were washed eight times with borate buffer.
  • the amount of spacer-bound protein was analysed by measuring the fluorescence intensity in the incubation plate at an excitation wavelength of 485 nm and an emission of 528 nm using a fluorescence reader (Synergy4, BioTek Instruments).
  • Known concentrations of labelled protein (6.25 - 500 ng) were used to prepare a calibration curve to convert fluorescence readings to amount of protein.
  • Sensors were manufactured as described in Example 2. Afterwards, the sensors were provided with a spacer layer as described in Example 2.
  • the spacer layer was made of Lipidure CM 5206 (NOF Corporation, Japan) or was made of AdaptTM (Biointeractions Ltd, Reading, England).
  • THP-1 cells were cultured in the presence of the sensors for 24 h at 37°C. The cells were then collected by centrifugation. The supernatant was used to determine the release of cytokines whereas the cell pellet was resuspended in PBS containing 1 % of bovine serum albumin (BSA) and used to analyse expression of the cell surface protein CD54 (also known as ICAM-1 , an inflammatory biomarker).
  • BSA bovine serum albumin
  • THP-1 cells were incubated with an anti-CD54 antibody conjugated with the fluorescent dye phycoerythrin (BD Bioscience). After incubation for 45 min at 4°C, cells were washed in PBS/ 1 % BSA and the mean fluorescence intensity (MFI) of 10000 cells was determined using a flow cyto meter (excitation wavelength 532 nm, emission wavelength 585 nm) (BD FACSArray, BD Bioscience). Compared to untreated THP-1 cells, incubation with sensors without coating resulted in increased relative CD54 expression (6-fold induction) as indicated by high MFI readings (Fig. 1 1 ). Incubating cells with sensors covered with spacer layers of CM 5206 or AdaptTM resulted in attenuation of CD54 expression by 45 % or 41 % respectively.
  • MFI mean fluorescence intensity
  • the supernatant was used to determine the amount of the cytokines Interleukin- 8 (IL-8) and "monocyte chemotactic protein- 1 " (MCP-1) using a bead-based immunoassay according to the manufacturer ' s instructions (Flex sets, BD Bioscience) and subsequent flow cytometric analysis (BD FACSArray, BD Bioscience). Data analysis was performed using FCAP array software v1 .0.1 (Soft flow Hungary Ltd.).
  • Adsorption of fibrinogen on the polymer surface and conformational changes in the protein might expose the MAC-1 binding site.
  • spacer layers made of AdaptTM or CM 5206 avoid protein deposition and exposure of structural motifs on surfaces (like sensors) and thereby minimise inflammatory reactions against implants.
  • Example 7 Limited haemolysis of sensors coated with a spacer layer
  • Sensors were manufactured as described in Example 2. Afterwards, the sensors were provided with a spacer layer as described in Example 2.
  • the spacer layer was made of Lipidure CM 5206 (NOF Corporation, Japan) or was made of AdaptTM (Biointeractions Ltd, Reading, England).
  • the haemolytic potential of sensors without spacer layer, sensors with spacer layer made of Lipidure CM 5206 or sensors with spacer layer made of AdaptTM was analysed. Therefore, sensors with a total surface area of 6 cm2 were incubated with red blood cells and then the lysis was determined by measuring the release of haemoglobin to the supernatant. Erythrocytes were isolated from fresh human blood by centrifugation (citrate was used to avoid coagulation).
  • erythrocytes were then washed with phosphate buffered saline (PBS) and thereafter diluted 1 :40 in PBS.
  • PBS phosphate buffered saline
  • the erythrocyte suspension was incubated with the sensors for 24 h at 37°C in the dark on a rotation platform (350 rpm). Afterwards, the cells were sedimented by centrifugation and the haemoglobin content of the supernatant was determined spectroscopically by measuring the absorption of the supernatant at a wavelength of 575 nm.

Abstract

La présente invention concerne un système d'électrodes pour mesurer la concentration d'un analyte en conditions in vivo, comprenant une électrode avec des molécules d'enzyme immobilisées et une barrière de diffusion améliorée qui régule la diffusion de l'analyte du fluide corporel entourant le système d'électrode aux molécules enzymatiques.
PCT/EP2013/056619 2011-03-28 2013-03-27 Membrane d'espacement amélioré pour un capteur enzymatique in vivo WO2013144255A1 (fr)

Priority Applications (17)

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DK13712795.7T DK2831263T3 (en) 2011-03-28 2013-03-27 IMPROVED SPACER MEMBRANE FOR AN ENZYMATIC IN-VIVO SENSOR
PL13712795T PL2831263T3 (pl) 2012-03-27 2013-03-27 Ulepszona membrana dystansująca do enzymatycznego czujnika in-vivo
ES13712795.7T ES2637811T3 (es) 2011-03-28 2013-03-27 Membrana espaciadora mejorada para un sensor enzimático in vivo
CA2867766A CA2867766C (fr) 2012-03-27 2013-03-27 Membrane d'espacement ameliore pour un capteur enzymatique in vivo
RU2014142913A RU2611038C2 (ru) 2012-03-27 2013-03-27 Улучшенная спейсерная мембрана для ферментного датчика in vivo
CN201380027747.XA CN104334740B (zh) 2012-03-27 2013-03-27 用于酶体内传感器的改进的隔离膜
EP13712795.7A EP2831263B1 (fr) 2012-03-27 2013-03-27 Membrane d'espacement amélioré pour un capteur enzymatique in vivo
PL17169444T PL3219807T3 (pl) 2012-03-27 2013-03-27 Ulepszona membrana dystansująca do enzymatycznego czujnika in-vivo
EP17169444.1A EP3219807B1 (fr) 2012-03-27 2013-03-27 Membrane d´espacement amélioré pour un capteur enzymatique in vivo
SI201330761T SI2831263T1 (sl) 2012-03-27 2013-03-27 Izboljšana razmična/vmesna membrana za encimski in-vivo senzor
BR112014021373-9A BR112014021373B1 (pt) 2011-03-28 2013-03-27 Sistema de eletrodo, sensor e uso de um copolímero hidrofílico
JP2015502344A JP6374860B2 (ja) 2012-03-27 2013-03-27 インビボ酵素センサー用の改良されたスペーサーメンブレン
IL234190A IL234190B (en) 2011-03-28 2014-08-19 An improved buffer membrane for an enzymatic in vivo sensor
ZA2014/06247A ZA201406247B (en) 2012-03-27 2014-08-25 Improved spacer membrane for an enzymatic in-vivo sensor
US14/486,402 US20150005605A1 (en) 2011-03-28 2014-09-15 Diffusion barriers and spacer membranes for enzymatic in-vivo sensors
HK15107281.6A HK1206793A1 (en) 2011-03-28 2015-07-30 Improved spacer membrane for an enzymatic in-vivo sensor
US18/194,967 US20230258595A1 (en) 2011-03-28 2023-04-03 Diffusion barriers and spacer membranes for enzymatic in-vivo sensors

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US11073495B2 (en) 2015-10-15 2021-07-27 Arkray, Inc. Biosensor and manufacturing method of biosensor
US11759132B2 (en) 2016-02-05 2023-09-19 Roche Diabetes Care, Inc. Medical device for detecting at least one analyte in a body fluid
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CN106957887A (zh) 2017-07-18
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