US20150005605A1

US20150005605A1 - Diffusion barriers and spacer membranes for enzymatic in-vivo sensors

Info

Publication number: US20150005605A1
Application number: US14/486,402
Authority: US
Inventors: Arnulf Staib; Marcel Thiele; Karl-Heinz Koelker; Ewald Rieger
Original assignee: Roche Diagnostics Operations Inc
Current assignee: Roche Diabetes Care Inc
Priority date: 2011-03-28
Filing date: 2014-09-15
Publication date: 2015-01-01
Also published as: US10422765B2; IL234190B; DK2697388T3; US20230258595A1; MX2013011041A; WO2012130841A1; CN103443286A; HK1206793A1; DK2831263T3; PL2697388T3; ES2637811T3; BR112014021373A2; ES2707347T3; EP2697388B1; AU2012234345A1; CN103443286B; EP2697388A1; HK1191984A1; BR112014021373B1; US20140018653A1

Abstract

An electrode system is disclosed for measuring a concentration or presence of an analyte under in-vivo conditions, where the electrode system includes at least one 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 diffusion barrier includes a hydrophilic polyurethane or a block copolymer having at least one hydrophilic block and at least one hydrophobic block. The electrode system also can include a spacer membrane that includes a hydrophilic copolymer of acrylic and/or methacrylic monomers.

Description

CROSS-REFERENCE TO RELATED APPLICATIONS

This patent application is a continuation of WO Patent Application No. PCT/EP2013/056619; filed 27 Mar. 2013, which claims priority to and the benefit of WO Patent Application No. PCT/EP2012/055406; filed 27 Mar. 2012, which claims priority to and the benefit of EP Patent Application No. 11160007.8, filed 28 Mar. 2011. Each patent application is incorporated herein by reference as if set forth in its entirety.

TECHNICAL FIELD

This disclosure relates generally to chemistry, engineering and medicine, and more particularly, it relates to electrode systems for measuring a concentration or a presence of an analyte under in-vivo conditions, where the electrode systems include an improved diffusion barrier that controls diffusion of the analyte from body fluid surrounding the electrode system to the enzyme molecules or an improved spacer layer that forms at least a portion of the outer layer of the electrode systems.

BACKGROUND

Sensors with implantable or insertable electrode systems facilitate measuring physiologically significant analytes such as, for example, lactate or glucose, in a patient's body. The working electrodes of these systems have electrically conductive enzyme layers in which enzyme molecules are bound and release charge carriers by catalytic conversion of analyte molecules. In the process, an electrical current is generated as a measuring signal whose amplitude correlates to the analyte concentration. Such electrode systems are known from, for example, WO Patent Application Publication Nos. 2007/147475 and 2010/028708.
The working electrodes of electrode systems often 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 systems to the enzyme molecules that are immobilized in the enzyme layer. For example, WO Patent Application Publication No. 2010/028708 discloses that the diffusion barrier of electrode systems can be a solid solution of at least two different polymers, such as acrylates. The polymers may be copolymers (e.g., copolymers of methyl methacrylate and hydroxyethyl methacrylate, or copolymers of butyl methacrylate and hydroxyethyl methacrylate).
Likewise, WO Patent Application Publication No. 2007/147475 discloses a diffusion barrier made from a polymer having a zwitterionic structure, such as poly(2-methacryloyloxyethyl phosphorylcholine-co-n-butylmethacrylate). The zwitterionic polymer may be mixed with another polymer such as, for example, polyurethane.
The use of polymer or copolymer mixtures, however, has drawbacks in that the preparation of the mixture and its application to the sensor is tedious and potentially problematic. Usually, the polymers to be mixed are individually dissolved, and the resulting solutions are thereafter mixed in the desired ratio. This may result in precipitation of one of the components and consequently in workability problems (e.g., in a spraying process). Increased difficulties occur when the mixture includes a polymer with ionic characteristics (i.e., when one of the polymers to be mixed includes a monomer having anionic or cationic groups). The presence of such charged groups has a strong effect on solubility, making it difficult to find a solvent suitable for both the charged polymer and an uncharged polymer.
WO Patent Application Publication No. 2006/058779 discloses an enzyme-based sensor with a combined diffusion and enzyme layer including at least one polymer material, and particles carrying an enzyme, where the particles are dispersed in the polymer material. The polymer also may include hydrophilic as well as hydrophobic polymer chain sequences. As such, the polymer may be a high or low water uptake polyether-polyurethane copolymer. The use of block copolymers having at least one hydrophilic block and at least one hydrophobic block as a diffusion layer is not disclosed.
EP Patent Application Publication No. 2 163 190 discloses an electrode system for measuring an analyte concentration in-vivo including a counter electrode with an electric conductor, and a working electrode with an electric conductor on which an enzyme layer including immobilized enzyme molecules is localized. A diffusion barrier controls diffusion of the analyte from surrounding body fluids or tissues to the enzyme molecules. The diffusion barrier may include hydrophilized polyurethanes obtained 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 including more than 50 mol-% hydrophilic monomers is not disclosed either.
For the foregoing reasons, there is a need for additional diffusion barriers, spacer membranes, electrode systems and sensors.

BRIEF SUMMARY

An inventive concept described herein includes block copolymer compositions of a single block copolymer having at least one hydrophilic block and at least one hydrophobic block, especially acrylic acid (acrylate) and/or methacrylic acid (methacrylate) units. The inventive concept is embodied in exemplary compositions, diffusion barriers, spacer membranes, electrode systems, sensors and methods of use thereof as described herein.
Advantageously, the block copolymer compositions provide desirable physico-chemical characteristics and that can be manufactured easily and incorporated into diffusion barriers and/or spacer membranes of an electrode system of an enzymatic in-vivo sensor.
In one aspect, block copolymer diffusion barriers are provided that include a single block copolymer having at least one hydrophilic block and at least one hydrophobic block (i.e., further polymers or copolymers are absent).
The hydrophilic and hydrophobic blocks are covalently linked to each other within a polymer molecule. The average molecular weight of the polymer (by weight) can be from about 20 kD to about 70 kD, from about 25 kD to about 60 kD, or from about 30 kD to about 50 kD.
The molar ratio of the hydrophilic to hydrophobic blocks in the block copolymer can be in the range from about 75% (hydrophilic):25% (hydrophobic) to about 25% (hydrophilic):75% (hydrophobic), from about 65% (hydrophilic):35% (hydrophobic) to about 35% (hydrophilic):65% (hydrophobic), or from about 60% (hydrophilic):40% (hydrophobic) to about 40% (hydrophilic):60% (hydrophobic).
A hydrophilic block of the block copolymer can be at least about 90%, at least about 95%, or completely of hydrophilic monomeric units. In some instances, it has a length from about 50-400 monomeric units, from about 50-200 monomeric units, from about 150-300 monomeric units, from about 100-150 monomeric units, or from about 200-250 monomeric units.
Likewise, a hydrophobic block of the copolymer can be at least 90%, at least about 95%, or completely of hydrophobic monomeric units. In some instances, it has a length from about 50-300 monomeric units, from about 50-200 monomeric units, from about 150-250 monomeric units, from about 80-150 monomeric units, or from about 170-200 monomeric units.
In some instances, the hydrophilic blocks and/or the hydrophobic blocks are (meth)acrylic-based units. In other instances, the hydrophilic blocks and the hydrophobic blocks both are (meth)acrylic-based monomeric units.
In another aspect, spacer membranes are provided that include a block copolymer, especially a hydrophilic copolymer of acrylic and/or methacrylic monomers. In some instances, the hydrophilic copolymer includes more than 50 mol-% hydrophilic monomers.
In another aspect, electrode systems are provided that include a block copolymer diffusion barrier as described herein. The electrode systems are for determining concentration or presence of an analyte under in-vivo conditions. In addition to the diffusion barrier, the electrode systems also include an electrode with immobilized enzyme molecules, where the diffusion barrier controls diffusion of the analyte from an exterior of the systems to the enzyme molecules.
The electrode systems are suitable for insertion or implantation into a body such as, for example, a mammalian body, especially a human body. The electrode systems also are adapted for determining concentration or presence of a desired analyte in a 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 systems are suitable for short-term application (e.g., about 3-14 days) or for long-term application (e.g., about 6-12 months). During the insertion or implantation period, the desired analyte may be determined by continuous or discontinuous measurements.
In some instance, the electrode systems can be part of an enzymatic, non-fluidic (ENF) sensor, where enzymatic conversion of the analyte is determined. Such a sensor includes a working electrode with immobilized enzyme molecules for converting the analyte and generating an electrical signal. The enzymes may be present in a layer covering the electrode. The sensor also can include redox mediators and/or electro-catalysts, as well as conductive particles and pore formers. In addition, the sensor can include a potentiostat and an amplifier for amplifying signals of the electrode system. In some instances, the sensor is for detecting the concentration or presence of glucose.
The area of the working electrode is the sensitive area of the sensor. As such, this sensitive area is provided with a block copolymer diffusion barrier as described herein. The diffusion barrier thus can be a cover layer (i.e., an enzyme-free layer) covering the enzyme layer. In some instances, the thickness of the diffusion barrier can be from about 2 μm to about 20 μm, from about 2 μm to about 15 μm, or from about 5 μm to about 20 μm. In other instances, the thickness of the diffusion barrier can be from about 5 μm to about 10 μm or from about 10 μm to about 15 μm (in dry state). Alternatively, it is feasible that diffusion-controlling particles can be incorporated into the enzyme layer to serve as a diffusion barrier. For example, pores of the enzyme layer can be filled with the block copolymer to control diffusion of analyte molecules.
The electrode system also can include a spacer membrane as described herein.
In view of the foregoing, methods are provided for using a block copolymer as described herein as part of an enzymatic electrode in an electrode system, where the block copolymer is incorporated into a diffusion barrier or a spacer membrane as described herein. The spacer membrane can be used to attenuate a foreign body reaction (FBR) against the enzymatic electrode when it is inserted or implanted into the body.
These and other advantages, effects, features and objects of the inventive concept will become better understood from the description that follows. In the description, reference is made to the accompanying drawings, which form a part hereof and in which there is shown by way of illustration, not limitation, embodiments of the inventive concept.

BRIEF DESCRIPTION OF THE DRAWINGS

The advantages, effects, features and objects other than those set forth above will become more readily apparent when consideration is given to the detailed description below. Such detailed description makes reference to the following drawings, wherein:

FIG. 1 shows an exemplary electrode system.

FIG. 2 shows a detailed view of FIG. 1.

FIG. 3 shows another detailed view of FIG. 1.

FIG. 4 shows a section along the section line CC of FIG. 2.

FIG. 5 shows sensitivity (with standard deviation) of four glucose sensors (at 10 mM glucose) provided with different block polymers (C, F, D, B) as diffusion barriers.

FIG. 6 shows sensor drift of four glucose sensors provided with different block copolymers (A, C, D, F) as diffusion barriers.

FIG. 7 shows conductivity of block copolymer A dependent on time (2 experiments).

FIG. 8 shows conductivity of block copolymer F dependent on time (3 experiments).

FIG. 9 shows conductivity of block copolymer H dependent on time for a layer thickness of 2.77 μm or 4.43 μm, respectively.

FIG. 10 shows fibrinogen adhesion to different spacer membrane polymers in-vitro (Adapt® and Eudragit® E100) with respect to an uncoated plate (Blank).

FIG. 11 shows expression of surface protein CD54 by THP-1 cells after incubation with sensors coated with a spacer membrane (Adapt™ and Lipidure® CM5206) or uncoated (control=untreated cells).

FIGS. 12 a-b show secretion of cytokine IL-8 and MCP-1, respectively, by THP-1 cells after incubation with sensors coated with a spacer membrane (Adapt™ and Lipidure® CM5206) or uncoated (control=untreated cells).

FIG. 13 shows 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 (polystyrene), and an additional layer of fibrinogen.

FIG. 14 shows the hemolysis after incubation with sensors coated with a spacer membrane (Adapt™ and Lipidure® CM5206) or uncoated in comparison to Adapt™ without sensor (negative control=incubation medium only; positive control=100% osmotic lysis).

While the inventive concept is susceptible to various modifications and alternative forms, exemplary embodiments thereof are shown by way of example in the drawings and are herein described in detail. It should be understood, however, that the description of exemplary embodiments that follows is not intended to limit the inventive concept to the particular forms disclosed, but on the contrary, the intention is to cover all advantages, effects, features and objects falling within the spirit and scope thereof as defined by the embodiments described herein and the claims below. Reference should therefore be made to the embodiments described herein and claims below for interpreting the scope of the inventive concept. As such, it should be noted that the embodiments described herein may have advantages, effects, features and objects useful in solving other problems.

DESCRIPTION OF EXEMPLARY EMBODIMENTS

The block copolymers, diffusion barriers, spacer membranes, electrode systems, sensors and methods now will be described more fully hereinafter with reference to the accompanying drawings, in which some, but not all embodiments of the inventive concept are shown. Indeed, the block copolymers, diffusion barriers, spacer membranes, electrode systems, sensors and methods may be embodied in many different forms and should not be construed as limited to the embodiments set forth herein; rather, these embodiments are provided so that this disclosure will satisfy applicable legal requirements.
Likewise, many modifications and other embodiments of the block copolymers, diffusion barriers, spacer membranes, electrode systems, sensors and methods described herein will come to mind to one of skill in the art to which the disclosure pertains having the benefit of the teachings presented in the foregoing descriptions and the associated drawings. Therefore, it is to be understood that the block copolymers, diffusion barriers, spacer membranes, electrode systems, sensors and methods are not to be limited to the specific embodiments disclosed and that modifications and other embodiments are intended to be included within the scope of the appended claims. Although specific terms are employed herein, they are used in a generic and descriptive sense only and not for purposes of limitation.
Unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of skill in the art to which the disclosure pertains. Although any methods and materials similar to or equivalent to those described herein can be used in the practice or testing of the block copolymers, diffusion barriers, spacer membranes, electrode systems, sensors and methods, the preferred methods and materials are described herein.
Moreover, reference to an element by the indefinite article “a” or “an” does not exclude the possibility that more than one element is present, unless the context clearly requires that there be one and only one element. The indefinite article “a” or “an” thus usually means “at least one.” Likewise, the terms “have,” “comprise” or “include” or any arbitrary grammatical variations thereof are used in a non-exclusive way. Thus, these terms may both refer to a situation in which, besides the feature introduced by these terms, no further features are present in the entity described in this context and to a situation in which one or more further features are present. For example, the expressions “A has B,” “A comprises B” and “A includes B” may both refer to a situation in which, besides B, no other element is present in A (i.e., a situation in which A solely and exclusively consists of B) or to a situation in which, besides B, one or more further elements are present in A, such as element C, elements C and D, or even further elements.
Overview
Exemplary diffusion barriers, spacer membranes, electrode systems, sensors and methods are provided and are based upon block copolymer compositions of a single block copolymer having at least one hydrophilic block and at least one hydrophobic block. Advantageously, the block copolymer compositions provide desirable physico-chemical characteristics and that can be manufactured easily and incorporated into diffusion barriers and/or spacer membranes of an electrode system of an enzymatic in-vivo sensor. Thus, electrode systems and sensors having diffusion barriers and/or spacer membranes as described herein can be used to attenuate FBR.
The diffusion barriers, spacer membranes, electrode systems and sensors are useful in a variety of applications. For example, the diffusion barriers, spacer membranes, electrode systems, sensors can be used to determine the concentration or presence of an analyte of interest in a sample such as, for example, determining the concentration or presence of glucose in a biological sample fluid.
The present inventive concept therefore provides block copolymers, diffusion barriers, spacer membranes, electrode systems, sensors and methods of using the same to attenuate sensor rejection.
Block Copolymers as Well as Diffusion Barriers and Spacer Membranes Incorporating the Same
Block copolymers of the inventive concept can include at least one hydrophilic block and at least one hydrophobic block. The hydrophilic monomeric units of the at least one hydrophilic block can be hydrophilic (meth)acryl esters (i.e., esters with a polar, such as, for example, OH, OCH₃or OC₂H₅group within the alcohol portion of the ester), hydrophilic (meth)acrylamides with an amide (NH₂) or an N-alkyl- or N,N-dialkylamide group, where the alkyl group includes 1-3 C-atoms and optionally hydrophilic groups such as OH, OCH₃or OC₂H₅, and suitable (meth)acrylic units having a charged (e.g., an anionic or cationic group such as, for example, acrylic acid (acrylate) or methacrylic acid (methacrylate)). It is contemplated that further combinations of monomeric units may be employed.
Other examples of monomeric units for the hydrophilic block include, but are not limited to, 2-hydroxyethyl acrylate, 2-hydroxyethyl methacrylate (HEMA), 2-methoxyethyl acrylate, 2-methoxyethyl methacrylate, 2-ethoxyethyl acrylate, 2-ethoxyethyl methacrylate, 2- or 3-hydroxypropyl acrylate, 2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA), 2- or 3-methoxypropyl acrylate, 2- or 3-methoxypropyl methacrylate, 2- or 3-ethoxypropyl acrylate, 2- or 3-ethoxypropyl methacrylate, 1- or 2-glycerol acrylate, 1- or 2-glycerol methacrylate, acrylamide, methacrylamide, an N-alkyl-or N,N-dialkyl acrylamide, and an N-alkyl- or N,N-dialkyl methylamide, where the alkyl includes 1-3 C-atoms such as methyl, ethyl or propyl, acrylic acid (acrylate), methacrylic acid (methacrylate) and combinations thereof.
In some instances, the hydrophilic monomers are HEMA and/or 2- or 3-HPMA. In other instances, the hydrophilic block includes 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.
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. In some instances, the hydrophilic block can be made from at least one non-ionic hydrophilic monomeric unit (e.g., as described above) and from at least one ionic hydrophilic monomeric unit, where the ionic monomeric unit is present in a molar amount of about 1 mol-% to about 20 mol-%. When the hydrophilic block includes the ionic monomeric unit, copolymerization can be with a hydrophilic monomer such as (meth)acrylamide, especially N,N-dialkyl acryl- or methacrylamide.
As used herein, “about” means within a statistically meaningful range of a value or values such as a stated concentration, length, molecular weight, pH, sequence identity, time frame, temperature or volume. Such a value or range can be within an order of magnitude, typically within 20%, more typically within 10%, and even more typically within 5% of a given value or range. The allowable variation encompassed by “about” will depend upon the particular system under study, and can be readily appreciated by one of skill in the art.
In some instances, the monomeric units for the hydrophilic block can be hydrophilic polyurethanes produced by polyaddition of an (poly)diisocyanat (e.g., 4-4-methylene-bis(cyclohexylisocyanate)) with a polyalcohol, such as a diol mixture. The components of the diol mixture can be polyalkylene glycols, such as polyethylene glycol (PEG) and polypropylene glycol (PPG) and aliphatic diols, such as ethylene glycol.
The hydrophilic polyurethane can be about 45-55 mol-% or even about 50 mol-% isocyanate and about 25-35 mol-% or even about 30 mol-% ethylene glycol. The degree of hydrophilization then can be adjusted by the ratio of PEG to PPG. In some instances, the polyurethane is about 2-3 mol-% or about 2.5 mol-% PEG, and about 17-18 mol-% or about 17.5 mol-% PPG.
To increase the hydrophilicity of the polyurethane, the proportion of PEG may be increased to about 4.5-5.5 mol-% or about 5 mol-% PEG to obtain an extremely hydrophilic polyurethane. The different hydrophilic variants of the polyurethanes also may be mixed to optimize the properties of the diffusion barrier.
Likewise, the hydrophobic monomeric units of the at least one hydrophobic block can be hydrophobic acrylic and/or methacrylic units, styrene-based monomeric units or combinations thereof. In some instances, the hydrophobic monomeric units can be hydrophobic (meth)acryl esters (e.g., esters having an alcohol portion with 1-3 C-atoms without hydrophilic group).
Other examples of monomeric units for the hydrophobic block include, but are not limited to, methyl acrylate, methyl methacrylate (MMA), ethyl acrylate, ethyl methacrylate (EMA), n- or i-propyl acrylate, n- or i-propyl methacrylate, n-butyl acrylate, n-butyl methacrylate (BUMA), neopentyl acrylate, neopentyl methacrylate, and combinations thereof.
In some instances, the hydrophobic block includes at least two different hydrophobic monomeric units that are present, for example, as a random copolymer. For example, the hydrophobic block can include MMA and BUMA, especially as a random copolymer of MMA and BUMA. The molar ratio between MMA and BUMA can be about 60% (MMA):40% (BUMA) to about 40% (MMA):60% (BUMA), or about 50% (MMA):50% (BUMA).
The glass transition (Tg) temperature of the hydrophobic block can be about 100° C. or less, about 90° C. or less, or about 80° C. or less (e.g., about 40-80° C.). As such, the hydrophobic block may be styrenic units (e.g., polystyrene having a Tg temperature of about 95° C.
The block copolymers described herein may be manufactured according to known methods such as, for example, the method disclosed in Boker et al. (2001) Macromolecules 34:7477-7488.
The block copolymers described herein can be incorporated into diffusion barriers and spacer membranes. As such, block copolymer-based diffusion barriers of the inventive concept include the block copolymers as described above. In addition to the block copolymer, the diffusion barriers also can include further components such as, for example, non-polymeric components, which may be dispersed and/or dissolved in the polymer. Examples of the further compounds include, but are not limited to, plasticizers, especially biocompatible plasticizers such as tri-(2-ethylhexyl) trimellitate and/or glycerol.
Likewise, block copolymer-based spacer membranes of the inventive concept include the block copolymers as described above. In particular, the spacer membranes include a hydrophilic copolymer of acrylic and/or methacrylic monomers, where the hydrophilic copolymer has more than about 50 mol-% hydrophilic monomers. Such spacer membranes can be used to attenuate a FBR against an enzymatic electrode, electrode system or sensor that includes the spacer membrane when it is inserted or implanted into the body.
The diffusion barriers/spacer membranes have a high effective diffusion coefficient (D_eff) for glucose, such as ≧10⁻¹⁰cm²/s, ≧5·10⁻¹⁰cm²/s, or ≧10⁻⁹cm²/s. Alternatively, D_effcan be up to 10⁻⁷or 10⁻⁸cm²/s at a temperature of about 37° C. and a pH of about 7.4. The D_effcan be determined as described in Example 4 according to the equation:
D _eff =SE _m /F·L _m·5182·10⁻⁸,
where SE_mis the sensitivity of the working electrode, F is the area of the working electrode, and L_mis the layer thickness of the diffusion barrier. SE_mand L_mmay be determined as described below in the Examples.
Block copolymer-based diffusion barriers and spacer membranes provide excellent physico-chemical characteristics as follows:
(i) permeability of the diffusion barriers/spacer membranes for the analyte to be determined;
(ii) permeability characteristics of the diffusion barriers/spacer membranes that are suitable for short-term behavior (wettability) and long-term behavior (sensor drift) of the electrode;
(iii) mechanical flexibility of the diffusion barriers/spacer membranes, which allows manufacture of in-vivo sensors with extended multiple electrodes; and/or
(iv) efficient incorporation of ionic groups into the diffusion barriers/spacer membranes (i.e., the density of cationic or anionic charges within the polymer can be efficiently adjusted), which is relevant for repulsion or attraction of charged analytes, and/or control of cell adhesion (e.g., of monocytes from the surrounding body fluid or tissue).
Electrode Systems and Sensors
Electrode systems of the inventive concept can include a block copolymer diffusion barrier and optionally a spacer membrane as described herein and at least one electrode such as a working electrode, where the at least one electrode includes immobilized enzyme molecules and an electrical conductor. In this manner, the diffusion barrier controls diffusion of the analyte from an exterior of the systems to the enzyme molecules, while the spacer membrane attenuates any FBR to the electrode system.
The enzyme immobilized in an enzyme layer on the at least one electrode is suitable for determining a concentration or a presence of a desired analyte. In general, the enzyme catalytically converts the analyte, thereby generating an electric signal detectable by the electric conductor of the at least one electrode such as a working electrode. Examples of enzymes include, but are not limited to, oxidases such as glucose oxidase or lactate oxidase, or a dehydrogenase such as glucose dehydrogenase or lactate dehydrogenase. Methods of immobilizing enzymes on electrodes are known in the art. See, e.g., WO Patent Application Publication No. 2007/147475.
The enzyme layer can be in the form of multiple fields that are arranged on the conductor of the working electrode at a distance of at least about 0.3 mm or at least about 0.5 mm from each other. In this manner, 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. By arranging the fields of the enzyme layer on the top of openings of an electrically insulating layer, the signal-to-noise ratio can be improved. Such an arrangement is disclosed in, for example, WO Patent Application Publication No. 2010/028708.
In addition to the enzyme, the enzyme layer can include an electro-catalyst or a redox mediator that transfers electrons to conductive components of the working electrode (e.g., graphite particles). Examples of electro-catalysts include, but are not limited to, metal oxides such as manganese dioxide, or organo-metallic compounds such as cobalt phthalo-cyanine. The redox mediator can degrade hydrogen peroxide, thereby counteracting depletion of oxygen in the surroundings of the working electrode.
The redox mediator may be covalently bound to the enzyme, thereby directly effecting electron transfer to the working electrode. Examples of redox mediators for direct electron transfer include, but are not limited to, prosthetic groups such as pyrrolo quinoline quinone (PQQ), flavine adenine dinucleotide (FAD) or other known prosthetic groups.
The electrode systems also can include a counter electrode with an electrical conductor. Likewise, the electrode systems can include a reference electrode capable of supplying a reference potential for the working electrode (e.g., an Ag/Ag—Cl reference electrode). Moreover, the electrode systems can include additional counter and/or working electrodes.
The electrode systems also can include an outer spacer membrane that covers at least the working electrode portion having the enzyme molecules and optionally also other portions (e.g., the counter electrode). If present, the spacer membrane also can cover the reference electrode. In some instances, the spacer membrane covers the entire implanted surface of the electrode system. In other instances, the spacer membrane covers the working electrode and optionally the counter electrode and the reference electrode, if present, in the form of a continuous layer.
The electrode systems may be part of a sensor in which the electrode system is connected to a potentiostat and an amplifier for amplifying signals of the electrode system. In some instances the sensor is an enzymatic non-fluidic (ENF) sensor, especially an electrochemical ENF sensor. In this manner, the electrodes of the electrode system may be arranged on a substrate that carries the potentiostat or may be attached to a circuit board that carries the potentiostat.
The electrode systems therefore are suitable for measuring the concentration or presence 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 such as, 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. Analytes of interest herein should be can be efficiently transferred between body fluid (e.g., blood or interstitial fluid) and tissue such as, for example, oxygen, carbon dioxide, sodium cations, chloride anions, glucose, urea, glycerol, lactate and pyruvate.
The block copolymers described above 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 such as, for example, an organic solvent like ether), which is applied to the prefabricated electrode system and dried thereon.
When the block copolymer is contacted with water, it shows a water uptake of about 15% to about 30% by weight (based on the polymer dry weight) at a temperature of about 37° C. and a pH of about 7.4 (aqueous phosphate buffer 10 mM KH₂PO₄, 10 mM NaH₂PO₄and 147 mM NaCl).
Referring now to FIG. 1, an exemplary electrode system is shown for insertion into body tissue of a human or animal such as, for example, into cutis or subcutaneous fatty tissue. A magnification of detail view A is shown in FIG. 2, a magnification of detail view B is shown in FIG. 3, and FIG. 4 shows a corresponding sectional view along the section line, CC, of FIG. 2.
The electrode system shown has a working electrode 1, a counter electrode 2 and a reference electrode 3. Electrical conductors of the electrodes 1 a, 2 a, 3 a are arranged in the form of metallic conductor paths and can be made of palladium or gold on a substrate 4. The substrate 4 can be a flexible plastic plate such as, for example, polyester. The substrate 4 typically is less than about 0.5 mm thick or is from about 100 μm to about 300 μm, and therefore is easy to bend so 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 can be at least about 1 cm in length or from about 2 cm to about 5 cm.
In the electrode system of FIG. 1, one part of the measuring facility, namely the head of the substrate, projects from the body of a patient during use. Alternatively, it is feasible just as well to implant the entire measuring facility and transmit measuring data in a wireless fashion to a receiver that is arranged outside the body.
As noted above, the working electrode 1 includes a enzyme layer 5 that contains immobilized enzyme molecules for catalytic conversion of the analyte. The enzyme layer 5 can be applied in, for example, 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 in, for example, WO Patent Application Publication No. 2007/147475.
The enzyme layer 5 need not be applied continuously on the conductor 1 a of the working electrode 1, but rather can be in the form of individual fields that are arranged at a distance from each other. The individual fields of the enzyme layer 5 in FIG. 1 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 FIG. 2. The conductor 2 a of the counter electrode 2 has a contour that follows the course of the conductor 1 a of the working electrode 1, resulting in an intercalating or interdigitated arrangement of working electrode 1 and counter electrode 2 with advantageously short current paths and low current density.
To increase its effective surface, the counter electrode 2 can be provided with a porous electrically conductive layer 6 that is situated in the form of individual fields on the conductor 2 a of the counter electrode 2. Like the enzyme layer 5 of the working electrode 1, 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 conductive layer 6 can have the same dimensions as the fields of the enzyme layer 5, although this is not obligatory. However, measures for increasing the surface of the counter electrode can just as well be foregone, and the counter electrode 2 can 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 membrane.
The reference electrode 3 can be arranged between the conductor 1 a of the working electrode 1 and the conductor 2 a of the counter electrode 2. The reference electrode shown in FIG. 3 includes a conductor 3 a on which a field 3 b of conductive silver/silver chloride paste is arranged.
As shown in FIG. 4, 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 counter electrode 2. Between the fields of enzyme layer 5, the conductor 1 a of the working electrode 1 can be covered with an electrically insulating layer 7, like the conductor 2 a of the counter electrode 2 between the fields of the conductive layers 6, to prevent interfering reactions that may otherwise be catalyzed by the metal of the conductor paths 1 a, 2 a. The fields of the enzyme layer 5 therefore are situated in openings of the insulation layer 7. Likewise, the fields of the conductive layer 6 of the counter electrode 2 may be placed on top of openings of the insulation layer 7.
The enzyme layer 5 can be 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 can be a single copolymer with alternating hydrophilic and hydrophobic blocks as described above.
The cover layer 8 can have a thickness from about 3 μm to about 30 μm, from about 5 μm to about 10 μm, or from about 10 μm to about 15 μm. 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. On the cover layer 8, a biocompatible membrane may be arranged as a spacer membrane 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 such 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 membrane 9 keep diffusing to the enzyme layer 5 of the working electrode 1 where they are converted. The spacer membrane 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. As shown, the spacer membrane 9 also cam cover the counter electrode 2 and the reference electrode 3.
In some instances, the spacer membrane 9 can be, for example, a dialysis membrane. As used herein, “dialysis membrane” means 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 can be selected so that analyte molecules can pass, while larger molecules are retained.
Alternatively, instead of a dialysis membrane, 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.
In some instances, the spacer membrane 9 is made from a copolymer of (meth)acrylates. In particular, the spacer membrane 9 can be a copolymer from at least 2 or 3 (meth)acrylates. As such, the spacer membrane 9 can include more than about 50 mol-%, at least about 60 mol-%, or at least about 70 mol-% hydrophilic monomer units (e.g. HEMA and/or 2-HPMA) and up to about 40 mol-% or up to about 30 mol-% hydrophobic monomer units (e.g., BUMA and/or MMA). The spacer membrane 9 may be a random or block copolymer. In certain instances, the spacer membrane 9 includes 2-HEMA and/or 2-HPMA as the hydrophilic monomers and includes MMA or BUMA as the hydrophobic monomers, where the amount of the hydrophilic monomers HEMA and/or HPMA is between about 80 mol-% to about 85 mol-%, and where the amount of hydrophobic monomers MMA and/or BUMA is between about 15 mol-% to about 20 mol-%. For example, the spacer membrane 9 can be made of the copolymer Adapt™ (Biolnteractions Ltd.; Reading, England), which includes BUMA as a hydrophobic moiety and 2-HEMA and 2-HPMA as hydrophilic moieties, where the amount of the 2-HEMA hydrophilic monomers is about 80 mol-%.
In any event, the spacer membrane 9 is highly permeable for the analyte (i.e., it does significantly lower the sensitivity per area of the working electrode such as, for example, about 20% or less, or about 5% or less) with a layer thickness of less than about 20 μm or even less than about 5 μm (e.g., from about 1 μm to about 3 μm).
The enzyme layer 5 of the electrode system can contain metal oxide particles, such as manganese dioxide particles, as catalytic redox mediator. Manganese dioxide catalytically converts hydrogen peroxide that is formed by enzymatic oxidation of glucose and other bioanalytes. During the degradation of hydrogen peroxide, the manganese dioxide particles transfer electrons to conductive components of the working electrode 1 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. Advantageously 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 spacer membrane 9 described herein may be used not only as an outer coating for a diffusion barrier 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. Thus, the spacer membrane can be arranged on the diffusion barrier; however, the spacer membrane also can be arranged directly on the enzyme layer. In this last context, the spacer membrane can act as the diffusion barrier itself and slow down diffusion of analyte molecules to the enzyme layer.
When the electrode system 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 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 about 50 wt-%, about 40 wt-%, or about 30 wt-% based on the total rate of the copolymer. Within the context of the present disclosure, the measurement of the relative water uptake can be performed by subjecting the dry copolymer to an excess of phosphate-buffer (pH 7.4) for about 48 hours at a temperature of about 37° C. The relative water uptake (WU %) is preferably determined according to the equation:
WU %=(m ₂ −m ₁)/m ₁×100,
where m₁and m₂represent, respectively, the mass of the dry copolymer and the copolymer after hydration according to the above measurement conditions.
It has been discovered that a 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 hours 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 Adapt™ advantageously increases the mechanical stability of the spacer membrane as described herein. In contrast, Lipidure® CM5206 shows a higher water uptake and swells to a hydrogel that is more fragile, easily deformable or shoved off, particularly when applied on an electrode system of an enzymatic in-vivo sensor.
Furthermore, during insertion and implementation of the electrode system, the spacer membrane is in direct contact with the tissue and/or the body fluid, like interstitial fluid or blood, containing biomolecules like proteins and cells. Advantageously, 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. In fact, reactions of the body against implanted material are known as “foreign body response” (FBR). During a 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, albumin, immunoglobulin, complement) to the surface of the former material (i.e., the implant). This protein coat presents binding sites to receptors on immune cells.
For example, fibrinogen contains a structural motif that binds to the monocyte receptor MAC-1. When 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. If the implant cannot be removed, a fibrous capsule is formed by connective tissue cells and proteins. This capsule, however, is a barrier for analytes to reach the sensor. Collectively, the events of the FBR are likely to interfere with the electrode system function in-vivo and with its life time.
Thus, an improved spacer membrane on an electrode-system of an enzymatic in-vivo sensor further provides a 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 (i.e., attenuates FBR). Thus, electrode systems for determining the concentration or presence of an analyte under in-vivo conditions includes an electrode with immobilized enzyme molecules and a diffusion barrier that controls diffusion of the analyte form the exterior of the electrode system to the enzyme molecules, as well as a spacer membrane that forms at least a portion of the outer layer of the electrode system, and where the spacer membrane includes a hydrophilic copolymer of acrylic and/or methacrylic monomers, especially more than about 50 mol-% hydrophilic monomers.
As described above, the spacer membrane 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. Examples 5 and 6 below show that an exemplary spacer membrane 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. Advantageously, the spacer membrane copolymer material does not activate immune cells itself. In Example 6, it could be demonstrated that an exemplary spacer membrane copolymer is able to attenuate the activation of immune cells by the implanted sensor. Moreover, advantageously, 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 also prevents hemolysis and complement activation by an implanted sensor. Thus, the spacer membrane 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 features of this embodiment, particularly with regard to the structure of the electrode system, the analyte and the enzyme molecules are as described herein. The diffusion barrier is as described herein; however, it also may have a different composition or may be absent. Accordingly, the diffusion barrier, when included, is a block copolymer having at least one hydrophilic block and at least one hydrophobic block as described herein.
Methods
Methods of the inventive concept include using a block copolymer having at least one hydrophilic block and at least one hydrophobic block as a diffusion barrier and/or spacer membrane for an enzymatic electrode. The block copolymer as described above (e.g., a single block copolymer). The diffusion barrier, the spacer membrane, the enzymatic electrode, electrode system and sensor are described above.

EXAMPLES

The inventive concept will be more fully understood upon consideration of the following non-limiting examples, which are offered for purposes of illustration, not limitation.

Example 1

Permeability of an Enzymatic Non-Fluidic Glucose Sensor with Distributed Electrodes for Transcutaneous Implantation Having a Diffusion Barrier Consisting of a Single Block Copolymer

Methods:
Sensors built on a prefabricated palladium strip conductor structure on a polyester substrate having a thickness of 250 μm. A working electrode (WE) and counter electrode (CE) were arranged distributedly as shown in, for example, FIGS. 1-2.
The fields of the CE were overprinted with carbon paste, and 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 AB; Sweden). The remaining paths of the strip conductor again were insulated. The reference electrode (RE) was a Ag/AgCl paste. The electrodes covered about 1 cm of the sensor shaft.
The WE-fields were coated with a block copolymer diffusion barrier of a HEMA block and a BUMA block. The thickness of the barrier was 7 μm.
Four sensor batches were produced, each provided with a specific block copolymer as a diffusion layer (see list below). All block copolymers were obtained from Polymer Source (Montreal, Canada) and are listed in the following Table 1.

TABLE 1

Block Polymer Diffusion Layer Characteristics.

	Molecular	Monomeric	Molecular
Name	Ratio/%	Units	Weight

Copolymer	BUMA/HEMA	HEMA	Copolymer [kD]
C	73/27	92	47
F	60/40	108	37
D	48/52	162	44
B	62/38	169	61

The respective block copolymers were dissolved in organic solvent (25% concentration) and the sensors were coated therewith. After drying by means of belt driers (2 min, 30° C. to 50° C.), the coated sensors were tested in-vitro in glucose solutions of different concentrations. Of each sensor batch, 10 sensors were measured as random sample. As a measure for the in-vitro sensitivity, the signal was calculated by the difference of the measured currents at 10 mM and 0 mM glucose concentration, which then was divided by 10 mM (cf. Example 4).
All sensors were operated at a polarization voltage of 350 mV versus Ag/AgCl, the measured temperature was kept constant at 37° C. The sensors used for this measurement series did not include a spacer membrane as described in WO Patent Application No. 2010/028708, which, however, did not make any difference in view of the tested signal level.
Results:
FIG. 5 shows sensor sensitivity with standard deviations for the four different diffusion layers. Concerning block copolymers C, D and F, there is a clear connection between in-vitro sensitivity and molar ratio of hydrophobic block when compared to hydrophilic block. At about identical total chain length of the copolymer, the sensitivity increases as the amount of hydrophilic block (HEMA) increases.
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 was 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 also may be seen in the gravimetrically determined water uptake of block copolymer B as compared to the other polymers. Polymer B has a water uptake of 10.6%±1.5% (weight percent referred to the polymer dry weight), whereas polymer C lies at 15.6%±0.0%, polymer F lies at 16.5±3.1%, and polymer D lies at 27%±1.7%.

Example 2

Mechanical Flexibility of a Diffusion Barrier of an ENF Glucose Sensor

Methods:
The sensors were manufactured as described in WO Patent Application Publication No. 2010/028708; however, the sensors included a diffusion barrier as described herein. It was assumed that the glass transition temperature (Tg) is a substitute parameter for the mechanic flexibility. Further, it was assumed that the Tg, 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 MMA-HEMA copolymer (produced by Polymer Source, Montreal). This polymer (called E) has a total molecular weight of 41 kD, and a molar ratio of MMA (hydrophobic amount) to HEMA is 60%:40%. The Tg of the hydrophobic block was 111° C., determined by DSC and a heating rate of 10° C./min.
Other sensors were provided with a diffusion barrier of a block copolymer (called A). The hydrophobic block of copolymer A contains MMA and BUMA at equal molar amounts in a randomized sequence. Again, the molar ratio of the hydrophobic part to the hydrophilic part was 60%:40%. The molecular weight was 36 kD. The Tg of the hydrophobic block decreased, due to the randomized sequence of MMA and BUMA (Tg about 45° C.), to 73° C.
Both diffusion barriers were generated from the respective solution (25%) of the copolymers in ether and dried as in Example 1. The thickness of the diffusion barriers was 7 μm. A spacer layer was applied subsequently via dip coating and dried 24 h at room temperature. The spacer layer was made of Lipidure® CM 5206.
Results:
After explanation from tissue, sensors having a copolymer E diffusion barrier showed sporadic cracks in the diffusion barrier. This is taken as an effect of the mechanic load. In contrast thereto, sensors having a copolymer A diffusion barrier did not show any cracks under identical load. This is due to a reduced Tg, which increases the mechanic stability of the copolymer. A physical mixture of two copolymers, as disclosed in WO Patent Application Publication No. 2010/028708, therefore is no longer required.

Example 3

Optimized Permeation Behavior of an ENF Glucose Sensor with Distributed Electrodes and a Diffusion Barrier

Methods:
A sensor was manufactured as described in Example 1, but with a spacer membrane on the total of the sensor shaft. Sensors with respective diffusion barrier were produced for copolymers A, C, D and F of Examples 1 and 2. For this purpose, a 24% etheric solution of the copolymer was generated. Each solution was applied onto a set of sensors (N=10) and then dried in a band drier. The resulting diffusion barriers had a thickness of 7 urn. Afterwards, 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. In-vitro measurements were carried out as in Example 1 but over a measuring period of 7 days. From the measured data, sensitivity drift was calculated over the respective measuring period for each sensor.
Results:
FIG. 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 6 h, the so-called startup phase—was excluded from the calculation.
For copolymers C, D and F having a hydrophobic block of BUMA, there was a positive drift (i.e., the sensitivity increased according to time). In contrast, copolymer A, with the hydrophobic block of a random copolymer of MMA and BUMA, led to a very low, slightly negative, drift.
These differences may be explained by the long-time permeability response of the respective diffusion barriers, which was measured in additional experiments. Palladium sensors without WE-paste, but with a defined active surface (i.e., also without an enzyme layer)—excluding the influence of its swelling behavior on the results—were coated with the above polymer solutions, and after drying, the thickness of the layer was measured. Subsequently, conductivity was measured in sodium- and chloride-containing buffer solution.
FIG. 7 shows that the conductivity of copolymer A remained nearly constant after a short startup phase. This was not the case for copolymer F, even under identical measurement conditions, as shown in FIG. 8. In this case, a long-term and strong permeability response of the copolymer F-based diffusion barrier was observed, which was practically independent of the layer thickness. For copolymer F—and also copolymers C and D (not shown)—with a hydrophobic block of BUMA, an increase of permeability results even over a long time period. When measured, this led to a continuous increase of sensitivity if the diffusion barrier was applied onto the sensor with distributed enzyme layer. This explains the observed positive sensor drift.
In contrast, a sensor having block copolymer A-based diffusion barrier showed a negligible drift, which was due to a very low permeability alteration in the conductivity measurement. Directly after starting measurement (until about 1 h afterwards), however, a strong increase of conductivity was observed in copolymer A. Here, a very fast startup was observed, which is terminated after about 1 hour. At this time, the diffusion layer was completely wetted and terminated its structural reorganization 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 reorganization, which is limited in time and amplitude, as compared to a copolymer having a Tg in the range of the ambient temperature.
In addition, sensors with copolymer A-based diffusion barriers showed a comparatively high sensitivity at the start of measurements when compared to sensors having a copolymer F-based diffusion barrier. This is to be expected due to the identical relative ratios between hydrophobic and hydrophilic blocks. The achieved sensitivity range of 1 nA/mM to 1.5 nA/mM (see, Example 1) is deemed ideal. This sensitivity likewise was obtained for sensors having a copolymer A-based diffusion barrier.
Regarding the sum of the three physico-chemical characteristics—permeability, mechanic stability and permeability response—an optimal sensor may be obtained with a diffusion barrier 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.

Example 4

Characterization of Block Copolymers

Methods:
A multiple field sensor (10 fields of working electrodes and counter electrodes, respectively) was produced and characterized in-vitro for continuously measurement of the glucose.
The sensor was provided with a diffusion barrier of a block copolymer having a hydrophobic block of random copolymerized MMA and BUMA and a hydrophilic block of HEMA. These polymers (specified G and H) were produced by Polymer Source (Montreal, Canada) and are more permeable than polymer A from Examples 1-3, which is included herein by reference. Table 2 below describes these copolymers.
Results:

TABLE 2

Block Polymer Diffusion Layer Characteristics.

Polymer

	G	H	A

Molecular weights Mn [kD]	23.5-b-29	21-b-20.5	21-b-15
Weight-% HEMA	55.2	49.4	41.6
Mol-% HEMA (stoichiometrically)	53.5	47.4	40
Mol-% HEMA (measured by ¹H, ¹³C	51	46	32.6
NMR)
Tg [° C.] hydrophobic block	65	68	86
HEMA monomeric units	223	157	115
MMA monomeric units	194	174	174

The molecular weights Mn of each block are separately indicated in 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 Tg temperatures of the hydrophobic blocks are within the desired range to guarantee mechanical flexibility.
The decisive parameter with regard to 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 a phosphate-buffered solution (pH 7.4) in nA/mM:
SE=[I(10 mM)−I(0 mM)]/10,
for each of the analyzed sensors.
From the individual measurement values (N=8), the mean sensitivity SE_mwas determined. The obtained sensitivity values were divided by the microscopically measured geometric total area F of all working electrode spots on the multi-field sensor. Thereby, a sensitivity density SE_m/F was obtained.
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·[I(20 mM)−I(0 mM)]/[I(10 mM)−I(0 mM)],
for each of the analyzed sensors.
From the individual measurement values, the mean linearity value and its standard deviation were determined (cf. Table 3).
Finally, 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. Therefrom, the effective diffusion coefficient D_oof the cover layer may be calculated:
D _eff =SE _m /F·L _m·5.182·10⁻⁸,
in cm²/s, wherein SE_mand L_mare the respective mean values for the sensitivity and the layer thickness, and F is the total area of all working electrode spots.
The sensor drift was calculated from repetitions of the glucose concentration stages over 7 days of in-vitro measurements. The results for polymer H showing a substantially constant conductivity are depicted in FIG. 9.

TABLE 3

Results of the Functional Characterization.

Polymer

	G	H

SE_m/F [nA/mM * mm²)]	1.85	1.25
Drift [% d]	−1.5 ± 0.2	0.3 ± 0.1
γ^{20 mM}[%]	88.2 ± 0.7	88.6 ± 0.3
layer thickness L _m[μm]	11.61	12.69
D_eff[cm²/s]	1.11305 * 10⁻⁹	8.22019 * 10⁻¹⁰

For the more hydrophilic polymer G (which is more permeable for glucose) the diffusion coefficient D_effalso was 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 D_effwas obtained (1.17·10⁻⁹cm²/s).

Example 5

Protein Binding to Spacer Layer Material

Methods:
To assess protein binding to spacer layer materials, ethanolic solutions of Adapt™ or Eudragit® E100 (Evonik Industries) were filled into an incubation plate (FluoroNunc Maxisorp; Thermo Scientific). 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 Alexa®488 (purchased from Invitrogen). After 4 h incubation, the fibrinogen solution was aspirated and the spacer layers were washed eight times with borate buffer. The amount of spacer-bound protein was analyzed by measuring 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 ng-500 ng) were used to prepare a calibration curve to convert fluorescence readings to amount of protein.
As expected, fibrinogen attached to the uncoated incubation plate (Blank) resulting in 390 ng of bound protein (FIG. 10). The plate coated with Eudragit® E100 showed reduced protein binding of 60 ng. Hardly any protein binding was detectable in the Adapt coated plate. The readings before incubation resulted from background fluorescence. These results clearly demonstrate that surfaces coated with the spacer materials, especially those coated with Adapt™, are well protected against fibrinogen adhesion.

Example 6

Cytokine Release by Cells after Contact with Spacer Layer

Methods:
Sensors were manufactured and provided with a spacer layer all as described in Example 2. The spacer layer was made of Lipidure® CM 5206 or was made of Adapt™
Sensors without spacer layers, sensors with spacer layers made of Lipidure® CM 5206, or sensors with spacer layers made of Adapt™ were incubated with monocytic THP-1 cells and induction of inflammatory markers was analyzed.
THP-1 cells were cultured in the presence of the sensors for 24 h at 37° C., and then were then collected by centrifugation. The supernatant was used to determine the release of cytokines, whereas the cell pellet was re-suspended 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). 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 cytometer (excitation wavelength 532 nm, emission wavelength 585 nm) (BD FACSArray, BD Bioscience).
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.).
To analyze the role of protein adsorption for activation of THP-1 cells, tissue culture plates were coated with Lipidure® CM 5206, Adapt™ or Eudragit® E100. The spacer was then incubated with human fibrinogen (Sigma-Aldrich). THP-1 cells were incubated with different spacers+fibrinogen layers. After 48 h incubation at 37° C. the cells were sedimented by centrifugation and the supernatant was analysed for IL-8 release. As a control, cells were grown in culture plates without spacer but coated with fibrinogen (polystyrene=culture plate material).
Results:
When 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. 11). Incubating cells with sensors covered with spacer layers of CM 5206 or Adapt™ resulted in attenuation of CD54 expression by 45 or 41% respectively.
In addition, when compared to untreated THP-1 cells, sensors without coating induced a strong release of IL-8 (49 vs.197 pg/ml; FIG. 12 a) and MCP-1(6 vs.48 pg/ml; FIG. 12 b). The release of IL-8 and MCP-1 was reduced when sensors were covered with spacer layers of Lipidure® CM 5206 or Adapt™. Specifically, sensors covered with Adapt™ induced the release of 100 pg/ml of IL-8 and 25 pg/ml of MCP-1. Likewise, sensors covered with Lipidure® CM 5206 resulted in secretion of 125 pg/ml of IL-8 and 18 pg/ml of MCP-1. Collectively, these data indicate that the spacer layers attenuate the induction of three well-known inflammation biomarker, namely CD54, IL-8 and MCP-1.
With respect to the role of protein adsorption for activation of THP-1 cells, FIG. 13 shows that the cells released 89 pg/ml of IL-8. Cells cultured on CM 5206+fibrinogen or on Adapt™+fibrinogen released 68 or 49 pg/ml, respectively. In contrast, cells grown on Eudragit E100+fibrinogen released 206 pg/ml of IL-8. Notably, cells grown on Eudragit E100 without fibrinogen-coating secreted only 59 pg/ml of IL-8. Adsorption of fibrinogen on the polymer surface and conformational changes in the protein might expose the MAC-1 binding site. THP-1 cells that are activated via binding to their MAC-1 receptor release cytokines, like IL-8, and thereby trigger an inflammatory response. Hence, spacer layers made of Adapt™ 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 Hemolysis of Sensors Coated with a Spacer Membrane

Methods:
Sensors were manufactured as described and were provided with a spacer membrane all as described in Example 2. The spacer layer was made of Lipidure® CM 5206 or was made of Adapt™
The hemolytic potential of sensors without spacer layers, sensors with spacer layers made of Lipidure® CM 5206 or sensors with spacer layers made of Adapt™ was analyzed. Therefore, sensors with a total surface area of 6 cm²were incubated with erythrocytes, and then the lysis was determined by measuring the release of hemoglobin to the supernatant.
Erythrocytes were isolated from fresh human blood by centrifugation (citrate was used to avoid coagulation). They were then washed with phosphate buffered saline (PBS) and thereafter diluted 1:40 in PBS. 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 hemoglobin content of the supernatant was determined spectroscopically by measuring absorption of the supernatant at a wavelength of 575 nm. The results are presented as lytic index in %, which is the release of hemoglobin in a sample divided by hemoglobin release in the positive control (=complete osmotic lysis of erythrocytes in distilled water). The results are shown in FIG. 14.
Results:
Sensors without a spacer layer significantly caused hemolysis as indicated by a high haemolytic index of 47.4%. Coating the sensors with a spacer layer of Lipidure® CM 5206 reduced the hemolytic potential of the sensors as demonstrated by a lytic index of 14.7%. Sensors coated with a spacer layer of Adapt™ marginally caused hemolysis resulting in a lytic index of 7.5%, which is in the range of the negative control (=erythrocytes in PBS incubated without any test material) or Adapt™ alone. These results indicated the protective function of the spacer layer to reduce hemolysis.
All of the patents, patent applications, patent application publications and other publications recited herein are hereby incorporated by reference as if set forth in their entirety.
The present inventive concept has been described in connection with what are presently considered to be the most practical and preferred embodiments. However, the inventive concept has been presented by way of illustration and is not intended to be limited to the disclosed embodiments. Accordingly, one of skill in the art will realize that the inventive concept is intended to encompass all modifications and alternative arrangements within the spirit and scope of the inventive concept as set forth in the appended claims.

Claims

The invention claimed is:

1. An electrode system for measuring a concentration or presence of an analyte under in-vivo conditions, the electrode system comprising:

an electrode having immobilized enzyme molecules thereupon;

an optional diffusion barrier that controls diffusion of the analyte from an exterior of the electrode system to the immobilized enzyme molecules; and

a spacer membrane that forms at least a portion of an outer layer of the electrode system, wherein the spacer membrane comprises a hydrophilic copolymer of acrylic and/or methacrylic monomers, and wherein the hydrophilic copolymer comprises more than 50 mol-% hydrophilic monomers.

2. The electrode system of claim 1, wherein the spacer membrane is a hydrophilic copolymer from at least two or three acrylic and/or methacrylic monomers.

3. The electrode system of claim 1, wherein the hydrophilic monomers are hydrophilic (meth)acrylesters with a polar group, hydrophilic (meth)acrylamides, (meth)acrylic acid, or combinations thereof.

4. The electrode system of claim 3, wherein the hydrophilic monomers are selected from the group consisting of 2-hydroxyethyl acrylate, 2-hydroxyethyl methacrylate (HEMA), 2-methoxyethyl acrylate, 2-methoxyethyl methacrylate, 2-ethoxyethyl acrylate, 2-ethoxyethyl methacrylate, 2- or 3-hydroxypropyl acrylate, 2- or 3-hydroxypropyl methacrylate (2- or 3-HPMA), 2- or 3-methoxypropyl acrylate, 2- or 3-methoxypropyl methacrylate, 2- or 3-ethoxypropyl acrylate, 2- or 3-ethoxypropyl methacrylate, 1- or 2-glycerol acrylate, 1- or 2-glycerol methacrylate, acrylamide, methacrylamide, an N-alkyl- or N,N-dialkyl acrylamide, an N-alkyl- or N,N-dialkyl methylamide, wherein the alkyl comprises 1-3 C-atoms, acrylic acid, methacrylic acid, and combinations thereof.

5. The electrode system of claim 4, wherein the hydrophilic monomers are HEMA, 2-HPMA, and combinations thereof.

6. The electrode system of claim 1, wherein the hydrophilic copolymer of the spacer membrane comprises at least about 60 mol-% or at least about 70 mol-% hydrophilic monomers.

7. The electrode system of claim 1, wherein the hydrophilic copolymer of the spacer membrane further comprises up to about 40 mol-% or up to about 30 mol-% hydrophobic monomers.

8. The electrode system of claim 7, wherein the hydrophobic monomers are selected from the group consisting of methyl acrylate, methyl methacrylate (MMA), ethyl acrylate, ethyl methacrylate (EMA), n- or i-propyl acrylate, n- or i-propyl methacrylate, n-butyl acrylate, n-butyl methacrylate (BUMA), neopentyl acrylate, neopentyl methacrylate, and combinations thereof.

9. The electrode system of claim 8, wherein the hydrophobic monomers are MMA, BUMA, and combinations thereof.

10. The electrode system of claim 8, wherein the hydrophobic monomers are MMA or BUMA, and the hydrophilic monomers are HEMA and/or 2-HPMA.

11. The electrode system of claim 8, wherein the spacer membrane comprises BUMA, 2-HEMA and 2-HPMA, and wherein the copolymer comprises about 80 mol-% of 2-HEMA monomers.

12. The electrode system of claim 1, wherein the hydrophilic copolymer is a random copolymer or a block copolymer.

13. The electrode system of claim 1, wherein the spacer membrane has a thickness that is less than about 20 μm, less than about 5 μm, or from about 1 μm to about 3 μm.

14. The electrode system of claim 1, wherein a relative water uptake of the hydrophilic copolymer does not exceed 50 wt-%, 40 wt-%, or 30 wt-%, based on the total weight of the copolymer.

15. The electrode system of claim 1, wherein the diffusion barrier comprises a hydrophilic polyurethane or a block copolymer having at least one hydrophilic block and at least one hydrophobic block.

16. The electrode system of claim 1, further comprising a counter electrode having a counter electrode electrical conductor, a working electrode having a working electrode electric conductor upon which the immobilized enzyme molecules are located, and wherein the diffusion barrier and the spacer membrane cover at least the working electrode and optionally the counter electrode.

17. A sensor that is insertable or implantable into a body comprising an electrode system of claim 1.

18. The sensor of claim 17, wherein the electrode system is configured for measuring glucose.

19. A spacer membrane for an electrode system comprising:

a hydrophilic copolymer of acrylic and/or methacrylic monomers, wherein the hydrophilic copolymer comprises more than about 50 mol-% hydrophilic monomers as the spacer membrane for the enzymatic electrode system.

20. The spacer membrane of claim 19, wherein the hydrophilic copolymer attenuates a foreign body reaction (FBR) against the enzymatic electrode system when compared to an enzymatic electrode system that lacks the spacer membrane.