Improved separator for an in-vivo sensorThe present application is a divisional application of the following applications: application date: year 2013, month 3, day 27; application No.: 201380027747.X (PCT/EP 2013/056619); the invention name is as follows: as above.
Technical Field
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 controlling the diffusion of the analyte from a body fluid surrounding the electrode system to the enzyme molecules.
Furthermore, 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 controlling the diffusion of the analyte from the outside of the electrode system to the enzyme molecules, and an improved spacer membrane forming at least a part of the outer layer of the electrode system.
Background
Sensors having implantable or insertable electrode systems are advantageous for measuring physiologically meaningful analytes in a patient, such as, for example, lactate or glucose. The working electrode of such systems has an electrically conductive enzyme layer in which enzyme molecules are incorporated which release charge carriers by catalytic conversion of analyte molecules. In this process, a current is generated as a measurement signal whose amplitude is related to the analyte concentration.
Such electrode systems are known, for example, from WO 2007/147475 and WO2010/028708, the contents of which are incorporated herein by reference.
The working electrode of the electrode system is provided with a diffusion barrier which controls the diffusion of the analyte to be determined from the body fluid or tissue surrounding the electrode system to the enzyme molecules immobilized in the enzyme layer. According to WO2010/028708, the diffusion barrier of the electrode system is a solid solution of at least two different polymers, preferably of acrylates. The polymer may be a copolymer, for example a copolymer of methyl methacrylate and hydroxyethyl methacrylate or a copolymer of butyl methacrylate and hydroxyethyl methacrylate.
WO 2007/147475 discloses diffusion barriers made of polymers having a zwitterionic structure. An example of such a polymer is poly (2-methacryloyloxyethyl phosphorylcholine-co-n-butyl methacrylate). The zwitterionic polymer may be mixed with another polymer, such as a polyurethane.
However, a disadvantage of using a mixture of polymers or copolymers is that the preparation of the mixture and its application to the sensor is cumbersome and may be problematic. Generally, the polymers to be mixed are dissolved separately and the resulting solutions are then mixed in the desired proportions. However, this may lead to precipitation of one of the components, leading to processability problems, for example during spraying. Greater difficulties arise when the mixture comprises polymers having ionic character, i.e. when one of the polymers to be mixed comprises a monomer having an anionic or cationic group. However, the presence of such charged groups has a strong influence on the solubility, making it difficult to find a solvent suitable for both charged and uncharged polymers.
WO 2006/058779 discloses an enzyme-based sensor having a combined diffusion and enzyme layer comprising at least one polymeric material, and particles carrying an enzyme, wherein the particles are dispersed in the polymeric material. The polymer may comprise hydrophilic as well as hydrophobic polymer chain sequences, for example the polymer may be a polyether-polyurethane copolymer with high or low water absorption. The use of block copolymers having at least one hydrophilic block and at least one hydrophobic block as diffusion layers is not disclosed.
EP- ｃA-2163190 describes an electrode system for measuring the concentration of an analyte in vivo, comprising ｃA counter electrode (counter electrode) with an electrical conductor, and ｃA working electrode with an electrical conductor on which an enzyme layer comprising immobilized enzyme molecules is arranged. The diffusion barrier controls the diffusion of analytes from the surrounding body fluid to the enzyme molecules. The diffusion barrier may comprise a hydrophilized polyurethane, which may be obtained by polycondensation of 4,4' -methylene-bis (cyclohexyl isocyanate) and a diol mixture, which may be polyethylene glycol and polypropylene glycol. The hydrophilic polyurethane layer may be covered with a separator, such as a copolymer of butyl methacrylate and 2-methacryloyloxyethyl-phosphorylcholine. The use of block copolymers having at least one hydrophilic block and at least one hydrophobic block as diffusion layers is not disclosed. Nor is there any disclosure of using a hydrophilic copolymer of (meth) acrylic monomers comprising more than 50 mol-% of hydrophilic monomers.
Summary of The Invention
It is an object of the present invention to provide a diffusion barrier on the electrode system of an in vivo sensor of an enzyme (enzyme), which provides the desired physicochemical properties, and which can be easily manufactured.
This object is achieved by providing 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. Preferably, the block is a (meth) acrylate polymer block.
The diffusion barrier based on block copolymers provides excellent physicochemical properties as follows:
(i) the permeability of the diffusion barrier to the analyte to be determined,
(ii) the permeability properties of the diffusion barrier, which are adapted to the short-term behavior (wettability) and the long-term behavior (sensor drift) of the electrode,
(iii) the mechanical flexibility of the diffusion barrier, which allows the fabrication of in vivo sensors with extended multiple electrodes;
(iv) the efficient introduction of ionic groups into the diffusion layer, i.e. the density of cationic or anionic charges within the polymer can be effectively modulated, which is related to the repulsion or attraction of charged analytes and/or the control of cell adhesion (e.g. to monocytes from surrounding body fluids or tissues).
The subject of the 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 controlling the diffusion of the analyte from the outside of the electrode system to the enzyme molecules, characterized in that the diffusion barrier comprises a block copolymer with at least one hydrophilic block and at least one hydrophobic block.
Preferably, 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. no further polymer or copolymer is present. More preferably, the diffusion barrier is composed of a single block copolymer having at least one hydrophilic block and at least one hydrophobic block.
The electrode system of the invention is suitable for insertion or implantation into the body, for example into the body of a mammal such as a human. The electrode system is suitable for measuring a desired analyte in a body fluid and/or body tissue, for example in an extracellular space (small gaps), in blood or in lymphatic vessels or in intercellular spaces.
The inserted or implanted electrode system is suitable for short term use, e.g. 3-14 days, or long term use, e.g. 6-12 months. During this insertion or implantation period, the desired analyte can be determined by continuous or discontinuous measurements.
The electrode system of the present invention is preferably part of an enzymatic non-fluid (ENF) sensor, wherein the enzymatic conversion of an analyte is determined. Preferably the sensor comprises a working electrode having an immobilized enzyme molecule for converting the analyte, the conversion resulting in the generation of an electrical signal. The enzyme may be present in a layer covering the electrode. In addition, redox mediators and/or electrocatalysts may be present as well as conductive particles and pore formers. Electrodes of this type are described, for example, in WO 2007/147475, the contents of which are incorporated herein by reference.
The area of the working electrode is the sensitive area of the sensor. The sensitive region is provided with a diffusion barrier which controls diffusion of the analyte from the outside, e.g. body fluids and/or tissues surrounding the electrode system, to the enzyme molecules. The diffusion barrier may for example be a cover layer covering the enzyme layer, i.e. an enzyme-free layer. However, it is also feasible to incorporate diffusion control particles into the enzyme layer to act as a diffusion barrier. For example, the pores of the enzyme layer may be filled with a polymer that controls the diffusion of analyte molecules. The thickness of the diffusion barrier is typically about 2-20 μm, e.g. about 2-15 μm, or about 5-20 μm, especially about 5-10 μm or about 10-15 μm (in the 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 blocks of alternating sequence, i.e. hydrophilic blocks linked to hydrophobic blocks. The hydrophilic and hydrophobic blocks are covalently linked to each other within the polymer molecule. The average molecular weight (by weight) of the polymer is generally from 20 to 70kD, particularly from 25 to 60kD, and more particularly from 30 to 50 kD. The molar ratio of hydrophilic to hydrophobic moieties in the block copolymer is typically about 75% (hydrophilic): 25% (hydrophobic) to about 25% (hydrophilic): 75% (hydrophobic), about 65% (hydrophilic): 35% (hydrophobic) to about 35% (hydrophilic): 65% (hydrophobic) or about 60% (hydrophilic): 40% (hydrophobic) to about 40% (hydrophilic): 60% (hydrophobicity).
The hydrophilic block of the block copolymer consists of at least 90%, at least 95% and in particular entirely of hydrophilic monomer units. It is typically 50-400, e.g., 50-200, or 150-300, especially 100-150, or 200-250 monomer molecules in length. The hydrophobic block of the copolymer consists of at least 90%, more particularly at least 95% and even more particularly entirely of hydrophobic monomer units. It is generally from 50 to 300, for example from 50 to 200 or 150 and 250, in particular from 80 to 150 or 170 and 200, monomer units in length.
The hydrophilic block and/or the hydrophobic block preferably consist of (meth) acrylic acid-based units. More preferably, both the hydrophilic block and the hydrophobic block are composed of (meth) acrylic acid-based monomer units.
The hydrophilic monomer units of the hydrophilic block are preferably selected from hydrophilic (meth) acrylates, i.e., having polar groups in the alcohol portion of the ester (i.e., OH, OCH)3Or OC2H5) Esters of (i) with amides (NH)2) Or a N-alkyl-or N, N-dialkylamide group, wherein the alkyl group contains 1 to 3C atoms and optionally a hydrophilic group, such as OH, OCH3Or OC2H5And suitable (meth) acrylic units having a charged group, for example an anionic or cationic group, such as acrylic acid (acrylate) or methacrylic acid (methacrylate). Furthermore, combinations of monomer units may be used.
Specific examples of preferred monomer units for the hydrophilic block are selected from:
2-hydroxyethyl acrylate, a mixture of 2-hydroxyethyl acrylate,
2-hydroxyethyl methacrylate (HEMA),
2-methoxy ethyl acrylate is prepared by reacting 2-methoxy ethyl acrylate,
2-methoxy ethyl methacrylate, and a mixture thereof,
2-ethoxyethyl acrylate, which is a mixture of acrylic acid and ethyl acrylate,
2-ethoxyethyl methacrylate, which is a mixture of ethyl 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-glyceride acrylate ester,
1-or 2-glyceride of methacrylic acid,
the reaction mixture of an acrylic amide and a water-soluble acrylic amide,
the reaction mixture of a methacrylic acid amide and a water-soluble acrylic acid amide,
n-alkyl-or N, N-dialkylacrylamides, and
n-alkyl-or N, N-dialkylmethylamides in which the alkyl group contains 1 to 3C atoms, such as methyl, ethyl or propyl,
acrylic acid (acrylic ester),
methacrylic acid (methacrylate) and combinations thereof.
Preferred hydrophilic monomers are 2-hydroxyethyl methacrylate (HEMA) and/or 2-or 3-hydroxypropyl methacrylate (2-or 3-HPMA). More preferably, the hydrophilic block is composed of at least two different hydrophilic monomeric units. For example, it may be a random copolymer of at least two different hydrophilic monomer units, such as HEMA and 2-HPMA.
To introduce ionic groups into the monomer, charged monomer units, such as acrylic acid (acrylate) and/or methacrylic acid (methacrylate), may be incorporated into the hydrophilic block. Thus, in one embodiment of the invention, the hydrophilic block may be made from at least one non-ionic hydrophilic monomeric unit (e.g., as described above) and at least one ionic hydrophilic monomeric unit, wherein the ionic monomeric unit is present in a molar amount of preferably 1 to 20 mol%. In the case where the hydrophilic block comprises ionic monomer units, such as acrylic acid or methacrylic acid, copolymerization with a hydrophilic monomer selected from (meth) acrylamides, in particular N, N-dialkylacrylamides or N, N-dialkylmethacrylamides, is preferred.
The hydrophobic monomer units of the hydrophobic block are preferably selected from hydrophobic acrylic and/or methacrylic acid units, styrene-based monomer units, or combinations thereof. Preferably, the hydrophobic monomer units are selected from hydrophobic (meth) acrylates, such as esters having an alcohol moiety containing 1-3C atoms but no hydrophilic groups. Specific examples of monomer units for the hydrophobic block are selected from:
the reaction mixture of methyl acrylate and methyl acrylate,
methyl Methacrylate (MMA),
the reaction mixture of ethyl acrylate and water is reacted,
ethyl Methacrylate (EMA) is used,
n-or iso-propyl acrylate is used,
n-or iso-propyl methacrylate, and,
n-butyl acrylate is used as the monomer,
n-butyl methacrylate (BUMA),
the acrylic acid ester of the neopentyl ester,
neopentyl methacrylate and combinations thereof.
The hydrophobic block preferably comprises at least two different hydrophobic monomer units, which are present, for example, as a random copolymer. In a preferred embodiment, the hydrophobic block comprises Methyl Methacrylate (MMA) and n-butyl methacrylate (BUMA). In a particularly preferred embodiment, 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), for example about 50% (MMA): 50% (BUMA). The hydrophobic block preferably has a glass transition temperature of 100 ℃ or less, 90 ℃ or less or 80 ℃ or less, for example about 40 to 80 ℃. In an alternative embodiment, the hydrophobic block may consist of styrene units, for example of polystyrene with a glass transition temperature of about 95 ℃.
The block copolymer used in the present invention can be manufactured according to known methods (nanoribbon ker et al, Macromolecules34(2001), 7477-.
The block copolymer may be applied to the electrode system by conventional techniques, for example by providing a solution of the block copolymer in a suitable solvent or solvent mixture (e.g. an organic solvent such as an ether), applying the solution to a pre-formed electrode system and drying thereon.
When the block copolymer is contacted with water, it is at a temperature of 37 ℃ and a pH of 7.4 (aqueous phosphate buffer 10 mMKH)2PO4，10mM NaH2PO4And 147mM NaCI) exhibit a water uptake (water uptake) of preferably about 15 wt% to 30 wt% (based on dry polymer weight).
In addition to the block copolymer, the diffusion barrier may also comprise further components, in particular non-polymeric components, which may be dispersed and/or dissolved in the polymer. These additional compounds include plasticizers, particularly biocompatible plasticizers such as tri- (2-ethylhexyl) trimellitate and/or glycerol.
The diffusion barrier of the present invention has a high effective diffusion coefficient for glucose DeffPreferably ≥ 10 at a temperature of 37 ℃ and a pH of 7.4-10cm2S, more preferably ≥ 5.10-10cm2S, and even more preferably ≥ 10-9cm2S, e.g. up to 10-7Or 10-8cm2And s. The effective diffusion coefficient is preferably determined as described in example 4 according to the following equation:
Deff=SEm/F·Lm·5182∙10-8
wherein SEmIs the sensitivity of the working electrode, F is the area of the working electrode, and LmIs the layer thickness of the diffusion barrier. SEmAnd LmCan be determined as described in the examples.
The electrode system of the present invention is adapted to measure the concentration of an analyte under in vivo conditions (i.e., when inserted or implanted in vivo). The analyte may be any molecule or ion present in a tissue or body fluid, such as oxygen, carbon dioxide, salts (cationic and/or anionic), fats or fat components, carbohydrates or carbohydrate components, proteins or protein components, or other types of biomolecules. It is particularly preferred to determine analytes such as oxygen, carbon dioxide, sodium cations, chloride anions, glucose, urea, glycerol, lactate and pyruvate which are efficiently transferred between a body fluid, e.g. blood, and a tissue.
The electrode system comprises an enzyme immobilized on an electrode. The enzyme is suitable for the determination of the desired analyte. Preferably, the enzyme is capable of catalytically converting the analyte and thereby producing an electrical signal that can be detected by the electrical conductor of the working electrode. The enzyme used for measuring the analyte is preferably an oxidase, such as glucose oxidase or lactate oxidase or dehydrogenase, such as glucose dehydrogenase or lactate dehydrogenase. In addition to the enzyme, the enzyme layer may also contain an electrocatalyst or redox mediator, which facilitates the transfer of electrons to the conductive components of the working electrode (e.g., graphite particles). Suitable electrocatalysts are metal oxides (such as manganese dioxide) or organometallic compounds (such as cobalt phthalocyanine). In a preferred embodiment, the redox mediator is capable of degrading hydrogen peroxide, thereby counteracting oxygen depletion around the working electrode. In a different embodiment, the 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 pyrroloquinoline quinone (PQQ), Flavin Adenine Dinucleotide (FAD) or other known prosthetic groups. Enzymes immobilized on electrodes are described, for example, in WO 2007/147475, the contents of which are incorporated herein by reference.
A preferred embodiment of the electrode system comprises a counter electrode having an electrical conductor and a working electrode having an electrical conductor on which the enzyme layer and the diffusion barrier are disposed. The enzyme layer is preferably designed in the form of a plurality of blocks (fields) which are arranged at a distance from one another, for example at least 0.3 mm or at least 0.5mm, on the conductor of the working electrode. The pieces of the working electrode may form a series of individual working electrodes. Between the blocks, the conductor of the working electrode may be covered by an insulating layer. By arranging the patch of enzyme layer on top of the opening of the electrically insulating layer, the signal to noise ratio can be improved. Such an arrangement is disclosed in WO2010/028708, the contents of which are incorporated herein by reference.
The electrode system of the invention may additionally comprise a reference electrode capable of supplying a reference potential to the working electrode, for example an Ag/Ag-Cl reference electrode. Furthermore, the electrode system of the invention may have further counter electrodes and/or working electrodes.
The electrode system may be part of a sensor, for example by being connected to a potentiostat and to an amplifier for amplifying the measurement signal of the electrode system. The sensor is preferably an enzymatic non-fluid (ENF) sensor, more preferably an electrochemical ENF sensor. The electrodes of the electrode system may be disposed on a substrate carrying the potentiostat or attached to a circuit board carrying the potentiostat.
Another subject of the invention relates 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 enzyme electrode. The block copolymer is preferably as described above, for example a single block copolymer. The diffusion barrier and the enzyme electrode are also preferably as described above.
Drawings
Further details and advantages of the invention are explained on the basis of exemplary embodiments and with reference to the drawings.
Fig. 1 shows an exemplary embodiment of an electrode system of the present invention.
Fig. 2 shows a partial enlarged view of fig. 1.
Fig. 3 shows another partial enlarged view of fig. 1.
Fig. 4 shows a section along section line CC of fig. 2.
FIG. 5 shows the sensitivity (and standard deviation) of four glucose sensors (in 10mM 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 polymers (A, C, D, F) as barrier layers.
FIG. 7 shows the conductivity dependence of the block copolymer A on time (2 experiments).
FIG. 8 shows the conductivity dependence of the block copolymer F on time (3 experiments).
FIG. 9 shows the conductivity of the block copolymer H as a function of time for a layer thickness of 2.77 μm or 4.43 μm, respectively.
FIG. 10 shows the adhesion of fibrinogen relative to an uncoated plate (blank) to different insulation film polymers (Adapt @ and Eudragit E100) in vitro.
Fig. 11 shows expression of the surface protein CD54 by THP-1 cells after incubation with sensors coated or uncoated (control = untreated cells) with barrier membranes (Adapt and Lipidure CM 5206).
FIGS. 12a and 12b show the secretion of cytokines IL-8 and MCP-1 by THP-1 cells after incubation with sensors coated or uncoated with isolating membranes (Adapt and Lipidure CM5206) (control = untreated cells), respectively.
FIG. 13 shows the secretion of cytokine IL-8 by THP-1 cells after incubation with tissue culture plates coated or uncoated (Polyst.) with Adapt @, Lipidure CM5206 and Eudragit E100 and having an additional fibrinogen layer.
FIG. 14 shows hemolysis (negative control = incubation medium only; positive control = 100% osmotic lysis) of the insulation polymer Adapt without sensor after incubation with insulation film (Adapt and Lipidure CM5206) coated or uncoated sensors compared to Adapt @.
Detailed description of the invention
Fig. 1 shows an exemplary embodiment of an electrode system for insertion into body tissue of a human or animal, for example into the dermis or subcutaneous adipose tissue. A partial enlarged view of a is shown in fig. 2, and a partial enlarged view of B is shown in fig. 3. Fig. 4 shows a corresponding sectional view along the sectional line CC in fig. 2.
The electrode system shown has a working electrode 1, a counter electrode 2 and a reference electrode 3. The electrical conductors 1a, 2a, 3a of the electrodes are arranged on the substrate 4 in the form of metallic conductor paths, preferably made of palladium or gold. In the exemplary embodiment shown, the substrate 4 is a flexible plastic sheet, for example made of polyester. The substrate 4 is less than 0.5mm thick, for example 100 and 300 microns thick, and is therefore easily bendable so that it can accommodate movement of the surrounding body tissue after its insertion. The base 4 has a narrow stem (narrow draft) for insertion into the tissue of the patient's body and a wide head for connection to an electronic system disposed outside the body. The stem of the substrate 4 is preferably at least 1cm long, in particular 2cm to 5 cm.
In the exemplary embodiment shown, a portion of the measuring device, i.e. the head of the base, protrudes from the patient during use. Alternatively, it is also possible to implant the entire test facility and to transmit the measurement data wirelessly to a receiver arranged outside the body.
The working electrode 1 carries an enzyme layer 5, which contains immobilized enzyme molecules for the catalytic conversion of the analyte. The enzyme layer 5 may be applied, for example, in the form of a solidified paste of carbon particles, a polymeric binder, a redox mediator or electrocatalyst, and enzyme molecules. Details of the production of this type of enzyme layer 5 are disclosed in, for example, WO 2007/147475, which is incorporated herein by reference in its entirety.
In the exemplary embodiment shown, the enzyme layer 5 is not applied continuously to the conductor 1a of the working electrode 1, but rather in the form of individual blocks arranged at a distance from one another. The individual pieces of the enzyme layer 5 are in a series arrangement in the exemplary embodiment shown.
The conductor 1a of the working electrode 1 has narrow points between the enzyme layer blocks, which are particularly clearly visible in fig. 2. The conductor 2a of the counter electrode 2 has a contour following the course (court) of the conductor 1a of the working electrode 1. This means that an embedded or crossed arrangement of working electrode 1 and counter electrode 2 with advantageously short current paths and low current densities is produced.
To increase its effective surface, the counter electrode 2 may have a porous conductive layer 6, which is located in the form of a single block on the conductor 2a of the counter electrode 2. Similar to the enzyme layer 5 of the working electrode 1, this layer 6 may be applied in the form of a solidified paste of carbon particles and a polymer binder. The block of layer 6 is preferably of the same size as the block of enzyme layer 5, although this is not essential. However, measures to increase the counter electrode surface are also foreseen and the counter electrode 2 can also be designed as a linear conductor path without any type of coating or with a coating made of said block copolymer and optionally spacers.
The reference electrode 3 is disposed between the conductor 1a of the working electrode 1 and the conductor 2a of the counter electrode 2. The reference electrode shown in figure 3 consists of a conductor 3a on which a block 3b of conductive silver/silver chloride paste is disposed.
Fig. 4 shows a schematic cross-sectional view along the cross-sectional line CC in fig. 2. The cross-section line CC runs between one of the enzyme layer patches 5 of the working electrode 1 and the patch of the conductive layer 6 of the counter electrode 2. Between the patches of enzyme layer 5, the conductor 1a of the working electrode 1 may be covered with an electrically insulating layer 7, like the conductor 2a of the counter electrode 2 between the patches of electrically conductive layer 6, to prevent metal-catalysed interfering reactions that might otherwise pass through the conductor paths 1a, 2 a. The pieces of the enzyme layer 5 are thus located in the openings of the insulating layer 7. Likewise, the patches of conductive layer 6 of the counter electrode 2 may also be placed over the openings of the insulating layer 7.
The enzyme layer 5 is covered by a cover layer 8, which provides a diffusion resistance for the analyte to be measured and thus acts as a diffusion barrier. The diffusion barrier 8 is composed of a single copolymer having alternating hydrophilic and hydrophobic blocks as described above.
An advantageous thickness of the cover layer 8 is, for example, 3 to 30 μm, in particular about 5 to 10 μm or about 10 to 15 μm. Because of its diffusion resistance, the cover layer 8 results in fewer analyte molecules per unit time reaching the enzyme layer 5. The cover layer 8 thus reduces the rate of conversion of analyte molecules and thereby counteracts depletion of analyte concentration around the working electrode.
The covering layer 8 extends substantially continuously over the entire area of the conductor 1a of the working electrode 1. On the cover layer 8, a biocompatible membrane may be provided as a spacer 9, which establishes a minimum distance between the enzyme layer 5 and the cells of the surrounding body tissue. This advantageously creates a reservoir for analyte molecules, from which analyte molecules can reach the respective enzyme layer block 5 in the event of a brief disturbance of the fluid exchange (fluidxchange) around the enzyme layer block 5. If the exchange of body fluids around the electrode system is temporarily limited or even prevented, the analyte molecules stored in the separator 9 continue to diffuse to the enzyme layer 5 of the working electrode 1, where they are converted. The spacers 9 thus cause a significant depletion of the analyte concentration and a corresponding distortion of the measurement result to occur only after a significantly longer period of time. In the exemplary embodiment shown, the membrane forming the separator 9 also covers the counter electrode 2 and the reference electrode 3.
The separation membrane 9 may for example be a dialysis membrane. In this context, a dialysis membrane is understood to be a membrane that is impermeable to molecules larger than the largest dimension. The dialysis membrane can be prefabricated in a separate manufacturing process and can then be applied during the manufacturing of the electrode system. The maximum size of the molecules that the dialysis membrane is permeable is chosen such that analyte molecules can pass through, while larger molecules are retained.
Alternatively, instead of a dialysis membrane, a coating made of a polymer with a high permeability to analytes and water, for example a coating based on polyurethane or acrylate, can be applied as a separation membrane 9 on the electrode system.
Preferably, the spacer is made of a copolymer of (meth) acrylic esters. Preferably the barrier film is a copolymer of at least 2 or 3 (meth) acrylates. More preferably the barrier film comprises more than 50mol%, at least 60mol% or at least 70mol% of hydrophilic monomeric units, such as HEMA and/or 2-HPMA, and at most 40mol% or at most 30mol% of hydrophobic units, such as BUMA and/or MMA. The separator may be a random or block copolymer. Particularly preferred separation membranes comprise MMA or BUMA as the hydrophobic moiety and 2-HEMA and/or 2-HPMA as the hydrophilic moiety. Preferably, the amount of hydrophilic monomers HEMA and/or HPMA is 80 to 85 mol-% and the amount of hydrophobic components MMA and/or BUMA is 15 to 20 mol-%.
A very preferred release film of the invention is made of copolymer Adapt @ (Biointeractions Ltd, Reading, England). Adapt @ comprises BUMA as the hydrophobic portion and 2-HEMA and 2-HPMA as the hydrophilic portion, wherein the amount of the 2-HEMA hydrophilic monomers is about 80 mol-%.
The separator is highly permeable to the analyte, i.e., it does significantly reduce the sensitivity of the working electrode per area, e.g., 20% or less, or 5% or less, and has a layer thickness of less than about 20 μm, preferably less than about 5 μm. Particularly preferred barrier film thicknesses are from about 1 to about 3 μm.
The enzyme layer 5 of the electrode system may comprise metal oxide particles, preferably manganese dioxide particles, as catalytic redox mediator. Manganese dioxide catalyzes the conversion of hydrogen peroxide, which is formed, for example, by the enzymatic oxidation of glucose and other biological analytes. During the degradation of hydrogen peroxide, the manganese dioxide particles transfer electrons to the conductive components of the working electrode 1, such as graphite particles in the enzyme layer 5. The catalytic degradation of hydrogen peroxide counteracts any reduction of the oxygen concentration in the enzyme layer 5. Advantageously, this enables the conversion of the analyte to be measured in the enzyme layer 5 to be unrestricted by the local oxygen concentration. The use of a catalytic redox mediator thus counteracts the distortion of the measurement results due to the low oxygen concentration. Another advantage of the catalytic redox mediator is that it prevents the production of cell damaging concentrations of hydrogen peroxide.
The preferred separator polymers described herein may be used as an outer coating for a diffusion barrier according to the present invention, but may also be used as an outer coating for electrode systems in general, in particular for electrode systems for measuring analyte concentrations under in vivo conditions, comprising an electrode with immobilized enzyme molecules and a diffusion barrier controlling diffusion of analyte from outside the electrode system to the enzyme molecules. Thus, the isolation membrane may be provided on the diffusion barrier, but the isolation membrane may also be provided directly on the enzyme layer. In this last case, the isolating membrane may also act as a diffusion barrier itself and slow down the diffusion of analyte molecules to the enzyme layer.
When the electrode system of the present invention is inserted or implanted in the body, the isolation diaphragm is the interface between the implanted sensor and the surrounding body fluids or tissues. Thus, the isolation diaphragm of the present invention must be mechanically robust so that it neither deforms nor moves away from the sensor when exposed to bodily fluids or tissue. For this reason, the water absorption of the separator copolymer and the accompanying swelling of the copolymer must be limited, although the copolymer has an inherent hydrophilicity.
Preferably, the relative water absorption of the separator copolymer should not exceed 50 wt%, preferably 40 wt%, more preferably 30 wt%, based on the total rate of the copolymer. In the present case, the relative water absorption was measured by subjecting the dried copolymer to an excess of phosphate buffer (pH7.4) at a temperature of 37 ℃ for 48 h. The relative water absorption (WU%) is preferably determined according to the following equation:
WU% = (m2-m1)/m1x 100,
wherein m is1And m2The mass of the dried copolymer and the mass of the hydrated copolymer according to the above measurement conditions are shown, respectively.
The inventors of the present invention determined that preferred insulation films made of the copolymer Adapt absorb 33. + -. 1.8 wt.% of phosphate buffer, pH7.4, relative to their own weight over a period of 48 h at 37 ℃. Under the same conditions, the membrane of the polymer Lipidure CM5206 (NOF Corporation, Japan) absorbed 157 ± 9.7 wt% of phosphate buffer relative to its own weight. The lower water absorption of the polymer advantageously increases the mechanical stability of the separator of the invention. In contrast, Lipidure CM5206 exhibits higher water absorption and swells into a more fragile, easily deformable or detachable hydrogel, particularly when applied to the electrode system of an in-vivo sensor.
Furthermore, during insertion and implementation of the electrode system, the spacer is in direct contact with tissue and/or body fluids, such as interstitial fluid or blood and the like (containing biomolecules such as proteins) and cells. Preferably, the barrier film must protect the inserted and implanted sensor from the tissue and/or body fluid environment and thus minimize the tissue reaction of the body with the implant. In fact, the body's response to the implant material is called the "exosome response" (FBR). With FBR, the body tries to destroy the implant or, if it is not possible, to create a capsule in order to separate it from the surrounding tissue (the foreign body granuloma). The first step of the FBR reaction is to bind proteins (e.g., fibrinogen, albumin, immunoglobulins, complement) to the surface of the former material (i.e., the implant). The protein coats the receptor which presents the binding site to the immune cell. 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 a binding site for MAC-1. Thus, immune cells, such as monocytes, are recruited to the implant and activated, secreting enzymes and free radicals to attack the implant. In addition, immune cells secrete soluble factors, i.e., cytokines, to recruit and activate other immune cells and thereby amplify the immune response. If the implant cannot be removed, fibrous capsules are formed from connective tissue cells and proteins. However, the capsule is a diffusion barrier for the analyte to reach the sensor. In summary, the events of the foreign body reaction as described above may interfere with the function of the electrode system in vivo and its lifetime.
Thus, the improved separator on the electrode system of the in-body sensor further provides for a reduction in tissue response to the implant and inhibits the formation of a capsule separating the sensor from surrounding tissue and body fluids.
It is therefore another object of the present invention to provide an electrode system for measuring the concentration of an analyte under in vivo conditions, comprising an electrode with an immobilized enzyme molecule and preferably a diffusion barrier controlling diffusion of the analyte from outside the electrode system to the enzyme molecule, characterized in that a separator film forms at least a part of the outer layer of the electrode system, wherein the separator film comprises a hydrophilic copolymer of acrylic and/or methacrylic monomers, wherein the polymer comprises more than 50mol% of hydrophilic monomers.
As mentioned above, the separator of the present invention does have limited protein binding capacity to protect the electrode system of the sensor from protein adsorption, which may trigger immune cell responses and may limit or interfere with its in vivo performance. Examples 5 and 6 show that the preferred separator membranes of the present invention provide weak binding to fibrinogen and prevent conformational changes in fibrinogen that would result in exposure of the MAC-1 binding motif to monocytes. Advantageously, the barrier membrane copolymer material does not activate the immune cells themselves. In example 6, it can be shown that the separator copolymers of the present invention are capable of attenuating immune cell activation by implanted sensors. Furthermore, advantageously, the barrier membrane is a biocompatible material, in particular compatible with body fluids, such as blood. Example 7 shows that the separator copolymer of the present invention can prevent hemolysis and complement activation of an implanted sensor. The barrier film of the invention therefore advantageously not only exhibits high mechanical stability, but also has optimal biocompatible properties, which is surprising because of the low water absorption when wetted.
Features of this embodiment, particularly in the structure of the electrode system, the analyte and the enzyme molecule, are as described herein. The diffusion barrier is preferably as described herein, but it may also have a different composition or may be absent. According to a preferred embodiment, the diffusion barrier preferably comprises a block copolymer having at least one hydrophilic block and at least one hydrophobic block as described above.
According to a further preferred embodiment, the diffusion barrier comprises a hydrophilic polyurethane. The hydrophilic polyurethanes used as diffusion membranes can be prepared by polyaddition (polyaddition) of (poly) diisocyanates, preferably 4-4-methylene-bis (cyclohexyl isocyanate), with polyols, preferably glycol mixtures.
The components of the glycol mixture are preferably polyalkylene glycols, such as polyethylene glycol (PEG) and polypropylene glycol (PPG), and aliphatic glycols, such as ethylene glycol. Preferably, the hydrophilic polyurethane comprises 45 to 55 mol-%, preferably 50 mol-%, isocyanate and 25 to 35 mol-%, preferably 30 mol-%, ethylene glycol. The degree of hydrophilization is then adjusted by the ratio of PEG to PPG. Preferably, the polyurethane comprises 2-3 mol-%, more preferably 2.5 mol-% PEG and 17-18 mol-%, preferably 17.5mol-% PPG. To increase the hydrophilicity of the polyurethane, the proportion of PEG can be increased, for example to 4.5-5.5 mol-%, preferably 5mol-% PEG, to obtain an extremely hydrophilic polyurethane. It is also possible to mix different hydrophilic variants of the polyurethane to optimize the properties of the diffusion barrier.
Preferred acrylic and methacrylic monomers for the release film copolymer are as described herein.
The hydrophilic monomeric units are preferably selected from hydrophilic (meth) acrylates, i.e., having polar groups in the alcohol portion of the ester (i.e., OH, OCH)3Or OC2H5) Esters of (i) with amides (NH)2) Or a N-alkyl-or N, N-dialkylamide group, wherein the alkyl group contains 1 to 3C atoms and optionally a hydrophilic group, such as OH, OCH3Or OC2H5And suitable (meth) acrylic units having a charged group, for example an anionic or cationic group, such as acrylic acid (acrylate) or methacrylic acid (methacrylate). Furthermore, combinations of monomer units may be used.
Specific examples of preferred monomer units for the hydrophilic block are selected from:
2-hydroxyethyl acrylate, a mixture of 2-hydroxyethyl acrylate,
2-hydroxyethyl methacrylate (HEMA),
2-methoxy ethyl acrylate is prepared by reacting 2-methoxy ethyl acrylate,
2-methoxy ethyl methacrylate, and a mixture thereof,
2-ethoxyethyl acrylate, which is a mixture of acrylic acid and ethyl acrylate,
2-ethoxyethyl methacrylate, which is a mixture of ethyl 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-glyceride acrylate ester,
1-or 2-glyceride of methacrylic acid,
the reaction mixture of an acrylic amide and a water-soluble acrylic amide,
the reaction mixture of a methacrylic acid amide and a water-soluble acrylic acid amide,
n-alkyl-or N, N-dialkylacrylamides, and
n-alkyl-or N, N-dialkylmethylamides in which the alkyl group contains 1 to 3C atoms, such as methyl, ethyl or propyl,
acrylic acid (acrylic ester),
methacrylic acid (methacrylate) and combinations thereof.
Preferred hydrophilic monomers are 2-hydroxyethyl methacrylate (HEMA) and/or 2-or 3-hydroxypropyl methacrylate (2-or 3-HPMA).
The hydrophobic monomer units are preferably selected from hydrophobic acrylic and/or methacrylic acid units or combinations thereof. Preferably the hydrophobic monomer units are selected from hydrophobic (meth) acrylates, for example esters having an alcohol moiety containing 1-3C atoms but no hydrophilic groups. Specific examples of monomer units for the hydrophobic block are selected from:
the reaction mixture of methyl acrylate and methyl acrylate,
methyl Methacrylate (MMA),
the reaction mixture of ethyl acrylate and water is reacted,
ethyl Methacrylate (EMA) is used,
n-or iso-propyl acrylate is used,
n-or iso-propyl methacrylate, and,
n-butyl acrylate is used as the monomer,
n-butyl methacrylate (BUMA),
the acrylic acid ester of the neopentyl ester,
neopentyl methacrylate and combinations thereof.
In a preferred embodiment, the hydrophobic block comprises Methyl Methacrylate (MMA) and n-butyl methacrylate (BUMA).
The outer separator film preferably covers at least the part of the working electrode comprising the enzyme molecule and optionally also other parts, such as the counter electrode. The separator also covers the reference electrode, if present. The isolating membrane preferably covers the entire implantation surface of the electrode system. The separator preferably covers the working electrode, optionally the counter electrode, and the reference electrode (if present in a continuous layer).
The electrode system comprising the improved isolation diaphragm of the present invention may be part of a sensor, for example by being connected to a potentiostat and an amplifier for amplifying the measurement signal of the electrode system. The sensor is preferably an enzymatic non-fluid (ENF) sensor, more preferably an electrochemical ENF sensor. The electrodes of the electrode system may be disposed on a substrate carrying the potentiostat or attached to a circuit board carrying the potentiostat. Preferably, the sensor is for measuring glucose.
A further subject of the invention relates to the use of a hydrophilic copolymer of acrylic and/or methacrylic monomers as a separator for an enzyme electrode, wherein the hydrophilic copolymer comprises more than 50 mol-% hydrophilic monomers. The hydrophilic copolymer is preferably as described above. Preferably, the isolation membrane serves to minimize an exosome reaction (FRB) against the enzyme electrode when it is inserted or implanted in vivo.
Examples
Example 1Enzymatic non-fluid (ENF) glucose sensor with distributed electrodes for percutaneous implantationAnd a diffusion layer composed of a single block copolymer.
The sensor was built onto a prefabricated palladium strip conductor structure on a polyester substrate 250 μm thick. The Working Electrodes (WE) and Counter Electrodes (CE) are arranged distributed (as shown in fig. 1-2).
The block of CE was overprinted with a carbon paste, insulating the rest of the strip conductor. Blocks of WE were overprinted with a mixture of cross-linked glucose oxidase (enzyme), conductive polymer paste and electrocatalyst, here manganese (IV) oxide (Technipur). The remaining paths of the strip conductors are insulated again. The Reference Electrode (RE) is made of an Ag/AgCl paste. These electrodes cover about 1cm of the sensor rod.
The WE block was coated with a block copolymer diffusion layer consisting of HEMA blocks and BUMA blocks. The thickness of this layer was 7 μm.
Four sensor batches were produced, each with a specific block copolymer as diffusion layer (see below list). All block copolymers were obtained from Polymer Source, Montreal and are listed in Table 1 below.
| Name (R) | Molecular ratio/%) | Monomer unit | Molecular 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 an organic solvent (25% concentration) and the sensors were coated therewith. After drying by means of a belt dryer (2min, 30-50 ℃), the coated sensors were tested in vitro in glucose solutions of different concentrations. In each sensor batch, 10 sensors were measured as random samples. As a measure for the in vitro sensitivity, the signal was calculated by the difference of the measured currents at 10mM and 0mM glucose concentration, which was then divided by 10mM (see example 4).
All sensors were operated at a polarization voltage of 350mV (vs. Ag/AgCl), keeping the measured temperature constant at 37 ℃. The sensors used in this measurement series do not contain the spacers described in WO2010/028708, but do not have any differences in view of the signal levels tested. Fig. 5 shows the sensor sensitivity at standard deviation for four different diffusion layers.
With respect to block copolymers C, D and F, there is a clear relationship between in vitro sensitivity and the molar ratio of hydrophobic block/hydrophilic block. At about the same total chain length of the copolymer, the sensitivity increases with increasing amount of hydrophilic block (HEMA).
The sensor with a diffusion layer of block copolymer B is an exception. Even though polymer B has a relative ratio of hydrophobic to hydrophilic amounts similar to polymer F, the sensitivity and thus the permeability to glucose is reduced. It can be said empirically that in the case of polymer B the total chain length, corresponding to the molecular weight of the copolymer molecules (total molecular weight), is so great that the permeability of the layer is reduced. This can also be seen in the gravimetric water absorption of the block copolymer B compared to the rest of the polymer. Polymer B has a water absorption of 10.6% + -1.5% (weight percent refers to dry polymer weight). Polymer C at 15.6% + -0.0%, polymer F at 16.5 + -3.1%, and polymer D at 27% + -1.7%.
Example 2Mechanism of diffusion layer of ENF glucose sensorFlexible fabricAnd (4) toughness.
The sensor was made as described in WO2010/028708 but with the diffusion layer of the present invention. The glass transition temperature (Tg) is assumed to be an alternative parameter for mechanical flexibility. Furthermore, given this glass transition temperature, which can be assigned to the hydrophobic block, determines the mechanical flexibility in vivo applications. It should be noted that a block copolymer may identify several tgs, which correspond to the number of blocks.
The sensor was coated with the same electrode paste as in example 1. Then, some sensors were coated with a copolymer selected from MMA-HEMA (produced by Polymer Source of Montreal). The total molecular weight of this polymer (designated E) was 41kD, the molar ratio of MMA (hydrophobic amount) to HEMA was 60%: 40 percent. The glass transition temperature of the hydrophobic block is 111 ℃ as determined by DSC and a heating rate of 10 ℃/min.
In addition, other sensors are provided with a block copolymer (referred to as a) diffusion layer of the present invention. The hydrophobic block of copolymer a comprises MMA and BUMA in equimolar amounts in random order. Likewise, the molar ratio of hydrophobic moieties to hydrophilic moieties is 60%: 40 percent. The molecular weight is 36 kD. The Tg of the hydrophobic block was reduced to 73 ℃ due to the random order of MMA and BUMA (Tg about 45 ℃).
Both diffusion layers were produced from respective solutions (25%) of the copolymers in ether and dried as in example 1. The thickness of the diffusion layer was 7 μm. The barrier layer was then applied via dip coating and dried at room temperature for 24 h. The separator was made of Lipidure CM5206 produced by NOF of Japan.
After removal from the tissue, the sensor with the copolymer E diffusion layer showed sporadic cracks in the diffusion layer. This is considered to be a function of mechanical load. In contrast, the sensor with the copolymer a diffusion layer did not show any cracks under the same load. This is apparently due to the decrease in Tg, which improves the mechanical stability of the copolymer. A physical mixture of two copolymers as disclosed in WO2010/028708 is no longer required.
Example 3Optimized ENF glucose sensor with distributed electrodes and diffusion layer according to the inventionOsmotic behavior.
The sensor was fabricated as described in example 1, but with an additional isolation layer over the entire sensor rod. For copolymers A, C, D and F of examples 1 and 2, sensors with respective diffusion layers were prepared. To this end, a 24% ether solution of the copolymer was produced. Each solution was applied to a set of sensors (N =10) and then dried in a belt dryer. A diffusion layer having a thickness of 7 μm was thus obtained.
Thereafter, these sensors were provided with the isolation layer described in example 2.
The sensors are connected to a measurement system on the sensor head which transfers the measurement data to a data store. In vitro measurements were performed as in example 1, but over a 7 day measurement period. From the measurement data, the sensitivity drift of each sensor is calculated over the respective measurement period. Fig. 6 shows the average of the in vitro drift values of the set for each sensor variant, i.e. the sensor with the diffusion layer variant. The calculation excludes the initial phase of the measurement (first 6h, the so-called start-up phase).
For all copolymers C, D and F with hydrophobic block BUMA, there is a positive drift, i.e. the sensitivity increases with time. In contrast, copolymer a with a hydrophobic block of a random copolymer of MMA and BUMA resulted in a very low slightly negative drift.
These differences can be explained by the long-term permeability response of the respective diffusion layers, which was measured in further experiments. The palladium sensor (without the WE paste, but with a defined active surface, i.e. also without the enzyme layer, excluding its swelling behaviour's effect on the results) was coated with the above polymer solution and the thickness of the layer was measured after drying. Subsequently, the conductivity was measured in a buffer solution containing sodium and chloride.
Figure 7 shows that the conductivity of copolymer a remains nearly constant after a short start-up period.
As can be seen from fig. 8, this is not the case for copolymer F, even under the same measurement conditions. In this case, a long-term and strong permeability response of the diffusion layer of copolymer F is observed, which is practically independent of the layer thickness. For the copolymer F with the BUMA hydrophobic block and the copolymers C and D (not shown), an increase in permeability is caused even in the long term. This results in a continuous increase in sensitivity if a diffusion layer is applied to the sensor carrying the distributed enzyme layer when measuring. This explains the observed positive sensor drift.
Vice versa, the sensor with block copolymer a shows negligible drift due to very low permeability changes in the conductivity measurements. But immediately after the start of the measurement (for about 1h thereafter), a drastic increase in conductivity was observed in copolymer a. Here, a very fast onset is observed, which terminates after about 1 hour. The diffusion layer is now fully wetted and its structural reorganization is terminated by water absorption. The degree of structural change is presumably dependent on Tg. It appears that the copolymer with an increased Tg passes through reorganization, which is limited in time and magnitude compared to a copolymer with a Tg in the ambient temperature range.
In addition, it must be acknowledged that the sensor with copolymer a exhibits a high sensitivity at the start of the measurement comparable to the sensor with the diffusion layer of copolymer F. This is expected due to the same relative ratio between the hydrophobic and hydrophilic blocks. The sensitivity range 1-1.5nA/mM achieved (see example 1) is considered ideal. This sensitivity is also obtained for sensors having a diffusion layer composed of copolymer a.
With respect to the sum of the three physico-chemical properties-permeability, mechanical stability and permeability response, an ideal sensor can preferably be obtained with a diffusion layer of a block copolymer having a hydrophobic block carrying at least two different randomly arranged hydrophobic monomer units, such as block copolymer a. None of the other block copolymers, whose hydrophobic block consists of only a single monomer unit, has achieved a quality comparable to that of copolymer a in all three parameters.
Example 4Characterization of the Block copolymer.
Multi-patch sensors for continuous glucose measurement were produced (10 patches for working and counter electrodes, respectively) and characterized in vitro.
The sensor has a diffusion layer composed of a block copolymer comprising a hydrophobic block of randomly copolymerized Methyl Methacrylate (MMA) and n-butyl methacrylate (BUMA) and a hydrophilic block of 2-hydroxyethyl methacrylate (HEMA). These polymers (designated G and H) have been produced by Polymer Source, Montreal, and have greater permeability than Polymer A of examples 1-3, which are incorporated herein by reference.
In table 2 below, these copolymers are described:
| polymer and method of making same | G | H | A |
| Molecular weight Mn [ kD ]] | 23.5-b-29 | 21-b-20.5 | 21-b-15 |
| Weight% HEMA | 55.2 | 49.4 | 41.6 |
| Mol% HEMA (stoichiometric) | 53.5 | 47.4 | 40 |
| Mol% HEMA (by)1H，13C NMR measurement) | 51 | 46 | 32.6 |
| Tg [℃]Hydrophobic block | 65 | 68 | 86 |
| HEMA monomer Unit | 223 | 157 | 115 |
| MMA monomer unit | 194 | 174 | 174 |
The molecular weight Mn of each block is shown in Table 2 above and represents an average value, since it is known that the polymer has a molecular chain length distribution in the vicinity of a specific median. This also applies to the derived quantities in table 2.
The glass transition temperature of the hydrophobic block is shown to be in a desired range to ensure mechanical flexibility.
A decisive parameter of the diffusion barrier for the permeability of the analyte is the sensitivity per area unit of the working electrode area (i.e. the geometrical area). For each sensor analyzed, the sensitivity SE was calculated from the current (I) measurements in phosphate buffer solution (pH7.4) at 10mM and at 0mM glucose concentration, in nA/mM:
SE = [I(10 mM) - I(0 mM)]/10
from the individual measurements (N =8), the average sensitivity SE was determinedm. The obtained sensitivity values were divided by the microscopically measured total geometric area F of all working electrode points on the multi-block sensor. Thereby obtaining a sensitivity density SEm/F。
The linearity Y of the in vitro function curve is an indication of the diffusion control functionality of the polymer coating on the working electrode. For each sensor analyzed, it was calculated from current measurements at 20mM, 10mM and 0mM glucose concentrations in%:
Y20mM= 50·[I(20mM) - I(0mM)]/[I(10mM) - I(0mM)]
from these individual measurements, the mean linearity value and its standard deviation were determined (see table 3).
Finally, for each polymer, the layer thickness L of the diffusion barrier of the sensor was determined by optical measurement. The corresponding average values for samples > 23 sensors with the same polymer were calculated. From this, the effective diffusion coefficient D of the cover layer can be calculatedeffIn units of cm2/s：
Deff=SEm/F·Lm·5.182∙10-8
Wherein SEmAnd LmIs the respective average of sensitivity and layer thickness, and F is the total area of all working electrode points.
Sensor drift was calculated from repeated glucose concentration phases in 7-day in vitro measurements. The results for polymer H, which exhibited substantially constant conductivity, are depicted in fig. 9.
Table 3 below shows the results of the functional characterization:
| polymer and method of making same | G | H |
| SEm/F [nA/mM*mm²)] | 1.85 | 1.25 |
| Drift [% d] | -1.5±0.2 | 0.3±0.1 |
| Y20mM[%] | 88.2±0.7 | 88.6±0.3 |
| Layer thickness Lm [µm] | 11.61 | 12.69 |
| Deff [cm²/s] | 1.11305*10-9 | 8.22019*10-10 |
For the more hydrophilic polymer G (which is more permeable to glucose), the diffusion coefficient was also determined by an alternative method, such as the permeation of glucose from the chamber with the glucose solution through the polymer membrane to the chamber with the glucose-free buffer. According to this method, similar diffusion coefficient values (1.17 · 10) were obtained-9cm2/s)。
Example 5Binding of proteins to the barrier material.
To assess binding of proteins to the barrier material, an ethanol solution of adapt (bio interactions Ltd, Reading, England) or Eudragit E100 (Evonik Industries) was 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 polymer was dried at 40 ℃ overnight. Thereafter, the release material is covered with a fibrinogen solution. This solution contained fibrinogen from human plasma conjugated to the fluorescent dye Alexa488 (from Invitrogen). After 4h incubation, the fibrinogen solution was aspirated and the isolation layer was washed eight times with borate buffer. The amount of spacer-bound protein was analyzed by measuring the fluorescence intensity in the incubation plate using a fluorescence reader (Synergy4, BioTek Instruments) at an excitation wavelength of 485 nm and an emission wavelength of 528 nm. A calibration curve was prepared using known concentrations of marker protein (6.25-500 ng) to convert the fluorescence reading to the amount of protein.
As expected, fibrinogen bound to uncoated incubation plates (blank), resulting in 390 ng of bound protein (fig. 10). The plate coated with Eudragit E100 showed 60 ng of reduced protein binding. Hardly any protein binding was detected in the AdaptTM coated plates. The pre-incubation readings were due to background fluorescence. These results clearly show that surfaces coated with spacer material, especially surfaces coated with AdaptTM, are well protected against fibrinogen adhesion.
Example 6Cytokine release from cells after contact with the barrier layer.
A sensor was fabricated as described in example 2. Thereafter, a sensor having an isolation layer was provided as described in example 2. The barrier layer was made of Lipidure CM5206 (NOF Corporation, Japan) or AdaptTM (biointerctions Ltd, Reading, England).
Sensors without a spacer layer, sensors with a spacer layer made of Lipidure CM5206, and sensors with a spacer layer made of adapt tm were incubated with mononuclear THP-1 cells and analyzed for induction of inflammatory markers.
THP-1 cells were cultured at 37 ℃ for 24h in the presence of a sensor. Cells were then collected by centrifugation. The supernatant was used to determine cytokine release, while the cell pellet was resuspended in PBS containing 1% Bovine Serum Albumin (BSA) and analyzed for expression of the cell surface protein CD54 (also known as ICAM-1, an inflammatory biomarker). THP-1 cells were incubated with anti-CD 54 antibody conjugated to the fluorescent dye phycoerythrin (BD Bioscience). After incubation at 4 ℃ for 45 min, 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) (BDFACSArray, BD Bioscience). Incubation with the sensor without the coating resulted in increased relative CD54 expression (6-fold induction) as shown by the high MFI readings compared to untreated THP-1 cells (fig. 11). Incubation of cells with sensors covered with CM5206 or AdaptTM isolation layers resulted in 45% or 41% attenuation of CD54 expression, respectively.
The supernatant was used to determine the amount of the cytokines interleukin-8 (IL-8) and "monocyte chemoattractant protein-1" (MCP-1) using a bead-based immunoassay according to the manufacturer's instructions (Flex series, BD Bioscience) and a subsequent flow cytometry analysis (BD FACSArray, BD Bioscience). Data analysis was performed using FCAP array software v1.0.1 (Softflow Hungary Ltd.).
The uncoated sensor induced a strong release of IL-8(49 vs.197 pg/ml) (FIG. 12a) and MCP-1(6 vs.48 pg/ml) (FIG. 12b) compared to untreated THP-1 cells. The release of IL-8 and MCP-1 is reduced when the sensor is covered with a separate layer of CM5206 or AdaptTM. The sensor covered with AdaptTM induced the release of 100 pg/ml IL-8 and 25 pg/ml MCP-1. The sensor covered with CM5206 resulted in the secretion of 125 pg/ml IL-8 and 18 pg/ml MCP-1.
Taken together, these data indicate that this isolation layer attenuates the induction of three well-known biomarkers of inflammation, namely CD54, IL-8, and MCP-1.
To analyze the effect of protein adsorption on activated THP-1 cells, tissue culture plates were coated with CM5206, AdaptTM or eudragit e 100. The spacer was then incubated with human fibrinogen (Sigma-Aldrich). THP-1 cells were incubated with different spacer + fibrinogen layers. After incubation at 37 ℃ for 48 h, the cells were pelleted by centrifugation and the supernatant was analyzed for IL-8 release. As a control, cells were grown in culture plates without spacers but coated with fibrinogen (Polyst = culture plate material). As shown in FIG. 13, these cells released 89 pg/ml IL-8. Cells cultured on CM5206 + fibrinogen or AdaptTM + fibrinogen released 68 or 49 pg/ml, respectively. In contrast, cells grown on Eudragit E100 + fibrinogen released 206 pg/ml IL-8. Notably, cells grown on Eudragit E100 without fibrinogen coating secreted only 59 pg/ml IL-8. Adsorption of fibrinogen on the polymer surface and conformational changes in the protein may expose the MAC-1 binding site. THP-1 cells, activated via binding to their MAC-1 receptor, release cytokines such as IL-8 and thereby trigger an inflammatory response. Thus, isolation layers made of AdaptTM or CM5206 avoid protein deposition and structural motif exposure on the surface (e.g., sensor) and thereby minimize the inflammatory response to the implant.
Example 7Limited hemolysis of a sensor coated with a barrier layer.
A sensor was fabricated as described in example 2. Thereafter, a sensor having an isolation layer was provided as described in example 2. The barrier layer was made of Lipidure CM5206 (NOF Corporation, Japan) or AdaptTM (biointerctions Ltd, Reading, England).
The hemolytic potential of sensors without spacer layer, sensors with spacer layer made of Lipidure CM5206, or sensors with spacer layer made of AdaptTM was analyzed. Thus, a sensor with a total surface area of 6 cm2 was incubated with red blood cells and then lysed by measuring the release of hemoglobin into the supernatant. Erythrocytes were separated from fresh human blood by centrifugation (citrate was used to avoid clotting). They were then washed with Phosphate Buffered Saline (PBS) and subsequently diluted 1:40 in PBS. The red blood cell suspension was incubated with the sensor at 37 ℃ on a rotating platform (350 rpm) in the dark for 24 h. Thereafter, the cells were sedimented by centrifugation, and the hemoglobin content of the supernatant was determined spectroscopically by measuring the absorption of the supernatant at a wavelength of 575 nm. The results are expressed as lysis index in%, which is the release of hemoglobin in the sample divided by the release of hemoglobin in the positive control (= complete osmotic lysis of red blood cells in distilled water). The results are shown in fig. 14.
The sensor without isolation layer caused significant hemolysis as indicated by a high hemolysis index of 47.4%. Coating the sensor with a spacer layer of Lipidure CM5206 reduced the hemolytic potential of the sensor as indicated by a lysis index of 14.7%. The sensor coated with a separate layer of adapt causes hemolysis slightly, resulting in a lysis index of 7.5%, which is in the range of negative controls (= red blood cells in PBS incubated without any test material) or adapt (tm) alone. These results indicate that the barrier layer reduces the protective function of hemolysis.