JP6374860B2 - Improved spacer membrane for in vivo enzyme sensors - Google Patents

Description

  The present invention relates to an analyte concentration under in vivo conditions comprising an electrode with an immobilized enzyme molecule and an improved diffusion barrier that controls the diffusion of the analyte from the body fluid around the electrode system to the enzyme molecule. It relates to an electrode system for measuring.

  Furthermore, the present invention provides an improvement that forms an electrode with immobilized enzyme molecules and optionally a diffusion barrier that controls the diffusion of analytes from outside the electrode system into the enzyme molecule and at least a portion of the outer layer of the electrode system. And an electrode system for measuring an analyte concentration under in vivo conditions.

  Sensors with an implantable or insertable electrode system facilitate the measurement of physiologically important analytes such as lactate or glucose in the patient's body. The working electrode of this type of system comprises a conductive enzyme layer in which the enzyme molecules are bound, which releases the charge carriers by catalytic conversion of the analyte molecules. In the process, a current is generated as a measurement signal, the magnitude of which correlates with the analyte concentration.

  Such electrode systems are known, for example, from WO 2007/147475 and WO 2010/028708, the contents of which are incorporated herein by reference.

  The working electrode of the electrode system is provided with a diffusion barrier that controls the diffusion of the analyte to be measured from body fluids or tissues around the electrode system to enzyme molecules immobilized in the enzyme layer. According to WO 2010/028708, the diffusion barrier of the electrode system is a solid solution of at least two different polymers, preferably acrylates. The polymer may be a copolymer, such as a copolymer of methyl methacrylate and hydroxyethyl methacrylate or a copolymer of butyl methacrylate and hydroxyethyl methacrylate.

  WO 2007/147475 discloses a diffusion barrier made from a polymer having a zwitterionic structure. An example of such a polymer is poly (2-methacryloyloxyethyl phosphorylcholine-co-n-butyl methacrylate). The zwitterionic polymer may be mixed with another polymer, such as polyurethane.

  However, the use of a polymer or copolymer mixture has the disadvantage that the preparation of the mixture and its application to the sensor is cumbersome and potentially problematic. Usually, the polymers to be mixed are dissolved individually and the resulting solution is then mixed in the desired ratio. However, this can lead to precipitation of one of the components and can result in processability problems, for example in the spraying process. The difficulty is increased when the mixture contains a polymer with ionic properties, i.e. one of the polymers to be mixed contains a monomer having an anionic or cationic group. The presence of such charged groups, however, strongly affects solubility and makes it difficult to find a suitable solvent for both charged and uncharged polymers.

  Patent Document 1 discloses that an enzyme-based sensor including an integrated diffusion layer and an enzyme layer containing at least one polymer material, and that the particles carry an enzyme, and the particles are dispersed in the polymer material. Is disclosed. The polymer may include hydrophilic and hydrophobic polymer chain sequences, for example, the polymer may be a high or low moisture uptake polyether-polyurethane copolymer. The use of a block copolymer having at least one hydrophilic block and at least one hydrophobic block as a diffusion layer is not disclosed.

  Patent Document 2 discloses an analysis comprising a counter electrode provided with a conductor, and a working electrode provided with a conductor, in which an enzyme layer containing enzyme molecules immobilized thereon is localized. An electrode system for the in vivo measurement of the concentration of an object is described. The diffusion barrier controls the diffusion of the analyte from surrounding body fluids to the enzyme molecules. The diffusion barrier may comprise a hydrophilicized polyurethane that may be obtained by condensation polymerization of 4,4′-methylene-bis- (cyclohexyl isocyanate) and a diol mixture (which may be polyethylene glycol and polypropylene glycol). . The hydrophilic polyurethane layer can be covered with a spacer, such as a copolymer of butyl methacrylate and 2-methacryloyloxyethyl-phosphoryl chlorin. The use of a block copolymer having at least one hydrophilic block and at least one hydrophobic block as a diffusion layer is not disclosed. The use of hydrophilic copolymers of (meth) acrylic acid monomers containing more than 50 mol-% hydrophilic monomer is also not disclosed.

International Publication No. 2006/058779 European Patent Application Publication No. 2163190

  It is an object of the present invention to provide a diffusion barrier in an electrode system of an enzymatic in vivo sensor that provides the desired physicochemical properties and 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 bound to each other. Preferably, the block is a (meth) acrylic acid polymer block.

Block copolymer based diffusion barriers provide very good physicochemical properties such as:
(I) The measured analyte permeability of the diffusion barrier (ii) The diffusion characteristics of the diffusion barrier that are suitable for the short-term behavior (wetting) and long-term behavior (sensor drift) of the electrode (iii) extended multiples The mechanical flexibility of the diffusion barrier, which enables the production of in vivo sensors with different electrodes (iv) the efficient incorporation of ionic groups into the diffusion layer, ie the cationic or anionic charge density in the polymer It can be efficiently tuned, which relates to the repulsion or affinity of the charged analyte and / or the control of the adhesion of cells such as monocytes from surrounding body fluids or tissues.

  The subject of the present invention is to measure the concentration of an analyte under in vivo conditions comprising an electrode with an immobilized enzyme molecule and a diffusion barrier that controls the diffusion of the analyte from outside the electrode system into the enzyme molecule An electrode system for use in which a diffusion barrier comprises a block copolymer having at least one hydrophilic block and at least one hydrophobic block.

  Preferably, the diffusion barrier comprises a single, ie only one, block copolymer having at least one hydrophilic block and at least one hydrophobic block, ie no further polymer or copolymer is present. More preferably, the diffusion barrier consists of a single block copolymer having at least one hydrophilic block and at least one hydrophobic block.

  The electrode system of the present invention is suitable for insertion or implantation into the body, for example into a mammalian body such as a human body. The electrode system is adapted to measure desired analytes in body fluids and / or body tissues, such as, for example, in the extracellular space (stroma), blood vessels or lymph vessels, or in interstitial spaces between cells.

  The inserted or implanted electrode system is suitable for short-term applications, such as 3-14 days, and long-term applications, such as 6-12 months. During the period of insertion or implantation, the desired analyte can be measured by continuous or non-continuous measurements.

  The electrode system of the present invention is preferably part of an enzymatic, non-fluidic (ENF) sensor, where the enzymatic conversion of the analyte is measured. Preferably, the sensor includes a working electrode on which an enzyme molecule for analyte conversion is immobilized, which results in the generation of an electrical signal. Enzymes can be present in the layer covering the electrode. In addition, redox mediators and / or electrocatalysts, conductive particles and pore formers may be present. Such electrodes are described, for example, in WO 2007/147475, the disclosure of which is incorporated herein by reference.

  The working electrode area is the sensitivity area of the sensor. This sensitive region is provided with a diffusion barrier that controls the diffusion of the analyte from the outside, eg, body fluids and / or tissues around the electrode system, to the enzyme molecules. The diffusion barrier can be, for example, a cover layer that covers the enzyme layer, ie, a layer that does not contain an enzyme. However, it is also possible that the diffusion control particles are incorporated in the enzyme layer so as to function as a diffusion barrier. For example, the pores of the enzyme layer can be filled with a polymer that controls the diffusion of analyte molecules. The thickness of the diffusion barrier is usually between about 2-20 μm, for example about 2-15 μm, or about 5-20 μm, in particular 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 an alternating sequence of blocks, ie a sequence in which hydrophilic blocks are bound to a hydrophobic block. The hydrophilic and hydrophobic blocks are covalently bound to each other within the polymer molecule. The (weight) average molecular weight of the polymer is usually 20-70 kD, in particular 25-60 kD and more particularly 30-50 kD. The molar ratio of hydrophilic portion to hydrophobic portion in the block copolymer is usually in the range of 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): in the range of 60% (hydrophobic).

  The hydrophilic block of the block copolymer consists of at least 90%, at least 95% and in particular completely hydrophilic monomer units. This usually has a monomer molecular length of 50 to 400, such as 50 to 200, or 150 to 300, in particular 100 to 150, or 200 to 250. The hydrophobic block of the copolymer consists of at least 90%, at least 95% and in particular completely hydrophobic monomer units. This usually has a monomer unit length of 50 to 300, such as 50 to 200, or 150 to 250, in particular 80 to 150, or 170 to 200.

  The hydrophilic block and / or the hydrophobic block are preferably composed of (meth) acrylic units. More preferably, both the hydrophilic block and the hydrophobic block are composed of (meth) acrylic monomer units.

The hydrophilic monomer unit of the hydrophilic block is preferably a hydrophilic (meth) acrylic ester, ie an ester having a polar group, ie an ester having an OH, OCH ₃ or OC ₂ H ₅ group in the alcohol part of the ester. , A hydrophilic (meth) acrylamide having an amide (NH ₂ ) or N-alkyl or N, N-dialkylamide group, wherein the alkyl group is 1-3 carbon atoms and optionally eg OH, (Meth) acrylamide containing hydrophilic groups such as OCH ₃ or OC ₂ H ₅ groups, and charged groups such as acrylic acid (acrylate) or methacrylic acid (methacrylate), such as anionic or cationic groups Is selected from suitable (meth) acrylic acid units having In addition, combinations of monomer units can be used.

Specific examples of preferred monomer units for the hydrophilic block are selected from:
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,
N-alkyl or N, N-dialkylacrylamide and N-alkyl or N, N-dialkylmethylamide, wherein the alkyl group contains 1 to 3 carbon atoms such as methyl, ethyl or propyl,
Acrylic acid (acrylate),
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 consists of at least two different hydrophilic monomer units. For example, it may be a random copolymer of at least two different hydrophilic monomer units such as HEMA and 2-HPMA.

  In order 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 certain embodiments of the present invention, the hydrophilic block may be made from at least one nonionic hydrophilic monomer unit (eg, those described above) and at least one ionic hydrophilic monomer unit, wherein The ionic monomer units are preferably present in a molar amount of 1 to 20 mol-%. When the hydrophilic block comprises ionic monomer units such as, for example, acrylic acid or methacrylic acid, it is particularly copolymerized with a hydrophilic monomer selected from the group of (meth) acrylamides such as N, N-dialkylacrylic or methacrylamide Is preferable.

The hydrophobic monomer units of the hydrophobic block are preferably selected from hydrophobic acrylic and / or methacrylic acid units, styrenic monomer units, or combinations thereof. Preferably, the hydrophobic monomer unit is selected from hydrophobic (meth) acrylic acid esters, such as, for example, esters having an alcohol moiety with 1 to 3 carbon atoms that do not have a hydrophilic group. Specific examples of monomer units for the hydrophobic block are selected from:
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.

  The hydrophobic block comprises at least two different hydrophobic monomer units, eg present 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 glass transition temperature of the hydrophobic block is preferably 100 ° C. or lower, 90 ° C. or lower, or 80 ° C. or lower, for example, about 40-80 ° C. In an alternative embodiment, the hydrophobic block may consist of styrene units, such as polystyrene having a glass transition temperature of about 95 ° C.

  The block copolymer used in the present invention can be produced according to a known method (Boeker et al., Macromolecules 34 (2001), 7477-7488).

  The block copolymer provides a solution of the block copolymer in a suitable solvent or solvent mixture, for example an organic solvent such as ether, which is applied to the ready-made electrode system and dried on it by conventional techniques. Can be applied to the electrode system.

When the block copolymer is contacted with water, preferably at a temperature of 37 ° C. and pH 7.4 (10 mM KH ₂ PO ₄ , 10 mM NaH ₂ PO ₄ , and 147 mM NaCl aqueous phosphate buffer), preferably about 15-30 wt. % Shows moisture uptake (based on dry weight of polymer).

  In addition to the block copolymer, the diffusion barrier may also include additional components, particularly non-polymeric components, which may be dispersed and / or dissolved in the polymer. These further components include, for example, plasticizers, in particular biocompatible plasticizers such as tri- (2-ethylhexyl) trimellitate and / or glycerol.

（ここで、ＳＥ_mは作用電極の感度であり、Ｆは作用電極の面積であり、および、Ｌ_mは拡散バリアの層厚である）
によって決定される。ＳＥ_mおよびＬ_mは実施例に記載されるように決定され得る。 The diffusion barrier of the present invention preferably has a temperature of 37 ° C. and pH 7.4, preferably ≧ 10 ⁻¹⁰ cm ² / s, more preferably ≧ 5 × 10 ⁻¹⁰ cm ² / s, and even more preferably Has a high effective diffusion coefficient D _eff for glucose which is ≧ 10 ⁻⁹ cm ² / s and for example up to 10 ⁻⁷ or 10 ⁻⁸ cm ² / s. The effective diffusion coefficient is preferably the following formula as described in Example 4:

(Where SE _m is the sensitivity of the working electrode, F is the area of the working electrode, and L _m is the layer thickness of the diffusion barrier)
Determined by. SE _m and L _m can be determined as described in the examples.

  The electrode system of the present invention is suitable for measuring the concentration of an analyte under in vivo conditions, i.e. when inserted or implanted in the body. An analyte can be any molecule or ion present in a tissue or body fluid, such as oxygen, carbon dioxide, salt (cation and / or anion), fat or fat component, carbohydrate or carbohydrate component, protein or protein component, or other Types of biomolecules and the like. Particularly preferably, analytes such as oxygen, carbon dioxide, sodium cation, chlorine anion, glucose, urea, glycerol, lactic acid and pyruvic acid are measured, which can be efficiently transferred between body fluids and tissues such as blood. .

  The electrode system includes an enzyme immobilized on the electrode. The enzyme is suitable for measuring the desired analyte. Preferably, the enzyme is capable of catalytically converting the analyte and thereby producing an electrical signal that is detectable by the conductor of the working electrode. The enzyme for measuring the analyte is preferably an oxidase such as glucose oxidase or lactate oxidase, or a dehydrogenase such as glucose dehydrogenase or lactate dehydrogenase. In addition to the enzyme, the enzyme layer may also include an electrocatalyst or redox mediator that aids the transfer of electrons to the conductive components of the working electrode, such as 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 has the ability to decompose hydrogen peroxide, thus offsetting the depletion of oxygen around the working electrode. In another embodiment, the redox mediator may be covalently bound to the enzyme, thereby effecting 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 with an electrical conductor and a working electrode with an electrical conductor on which an enzyme layer and a diffusion barrier are disposed. The enzyme layer is preferably designed in the form of a plurality of fields arranged on the conductor of the working electrode, for example at a distance such as at least 0.3 mm or at least 0.5 mm from each other. The individual fields of the working electrode may form a series of individual working electrodes. During these fields, the conductor of the working electrode may be covered by an insulating layer. By placing the field of the enzyme layer over the opening of the electrically insulating layer, the signal to noise ratio can be improved. Such an arrangement is disclosed in WO 2010/028708, the contents of which are incorporated herein by reference.

  The electrode system of the present invention additionally comprises a reference electrode that can provide a reference potential for the working electrode, for example an Ag / Ag-Cl reference electrode. Furthermore, the electrode system of the present invention may have an additional counter and / or working electrode.

  The electrode system may be made part of the sensor, for example by being coupled to a potentiostat and an amplifier for amplification of the measurement signal of the electrode system. The sensor is preferably an enzymatic, non-fluidic (ENF) sensor, more preferably an electrochemical ENF sensor. The electrodes of the electrode system may be disposed on a substrate carrying a potentiostat or may be bonded to a circuit board carrying a potentiostat.

  A further subject matter of the present 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 enzymatic electrode. The block copolymer is preferably, for example, a single block copolymer, as described above. The diffusion barrier and enzymatic electrode are also preferably as described above.

  Further details and advantages of the invention will be explained on the basis of exemplary embodiments with reference to the attached drawings.

2 illustrates an exemplary embodiment of the electrode system of the present invention. FIG. 2 is a detailed view of FIG. 1. FIG. 3 is another detailed view of FIG. 1. FIG. 3 shows a cross-sectional view along the cross-sectional line CC of FIG. 2. The sensitivity (with standard deviation) of four glucose sensors provided with different block polymers (C, F, D, B) as barrier layers is shown. Fig. 4 shows the sensor drift of four glucose sensors provided with different block polymers (A, C, D, F) as barrier layers. The time-dependent conductivity of block copolymer A is shown (two experiments). The time-dependent conductivity of block copolymer F is shown (three experiments). Figure 2 shows the time-dependent conductivity of block copolymer H with a layer thickness of 2.77 µm or 4.43 µm, respectively. Figure 2 shows in vitro fibrinogen adhesion of various spacer membrane polymers (Adapt® and Eudragit® E100) to an uncoated plate (control). Surface protein CD54 by THP-1 cells after incubation with spacer membrane (Adapt® and Lipidure® CM5206) or uncoated (control = untreated cells) sensor Expression is shown. Cytokine IL- by THP-1 cells after incubation with sensors coated with spacer membranes (Adapt® and Lipidure® CM5206) or uncoated (control = untreated cells) sensor 8 secretions are shown. MCP-1 by THP-1 cells after incubation with sensors coated with spacer membranes (Adapt® and Lipidure® CM5206) or uncoated (control = untreated cells) sensor Showing secretion. By THP-1 cells after incubation with tissue membrane plates coated or uncoated (Polyst.) With spacer membranes (Adapt®, Lipidure® CM5206 and Eudragit® E100) , Shows secretion of cytokine IL-8. Spacer membranes (Adapt® and Lipidure®) compared to spacer polymer Adapt® without sensor (negative control = incubation medium only; positive control = 100% osmotic lysis) Shows hemolysis after incubation with CM5206) coated or uncoated sensor.

  FIG. 1 illustrates an exemplary embodiment of an electrode system for insertion into human or animal body tissue, such as skin or subcutaneous adipose tissue. An enlarged detail view of A is shown in FIG. 2, and an enlarged detail view of B is shown in FIG. FIG. 4 shows a corresponding cross-sectional view along the cross-sectional line CC of FIG.

  The electrode system shown comprises a working electrode 1, a counter electrode 2 and a reference electrode 3. The electrode conductors 1a, 2a, 3a are arranged on the substrate 4 in the form of metal conductor paths, preferably made of palladium or gold. In the exemplary embodiment shown, the substrate 4 is a flexible plastic plate, for example made from polyester. The substrate 4 is less than 0.5 mm, for example 100-300 micrometers thick, and is therefore easy to bend so that it can adapt to the movement of surrounding body tissue after its insertion. The substrate 4 comprises a narrow shaft for insertion into the patient's body tissue and a wide upper end for connection to an electrical system located outside the body. The shaft part of the substrate 4 is preferably at least 1 cm long, in particular 2 cm to 5 cm long.

  In the exemplary embodiment shown, a portion of the measurement mechanism, i.e., the top edge of the substrate, protrudes from the patient's body during use. Alternatively, however, it is also possible to embed the entire measurement mechanism and to transmit the measurement data in a wireless format to a receiver that is located outside the body.

  The working electrode 1 carries an enzyme layer 5 containing immobilized enzyme molecules for catalytic conversion of the analyte. The enzyme layer 5 can be applied in the form of, for example, a hardened paste of carbon particles, a polymer binder, a redox mediator or electrocatalyst, and enzyme molecules. Details of the production of this type of enzyme layer 5 are disclosed, for example, in WO 2007/147475, which is incorporated herein by reference.

  In the exemplary embodiment shown, the enzyme layer 5 is not applied continuously to the conductor 1a of the working electrode 1, but rather is applied in the form of individual fields arranged at a distance from one another. The The individual fields of the enzyme layer 5 in the exemplary embodiment shown are arranged as a series.

  The conductor 1a of the working electrode 1 has a narrow portion between the enzyme layer fields, as is particularly well shown in FIG. The conductor 2 a of the counter electrode 2 has an outer shape along the path of the conductor 1 a of the working electrode 1. This measure results in an arrangement in which the working electrode 1 and the counter electrode 2 are intercalated or combined with each other, advantageously with a short current path and low current density.

  In order to increase its effective surface, the counter electrode 2 can be provided with a porous conductive layer 6 located in the form of individual fields on the conductor 2a of the counter electrode 2. Similar to the enzyme layer 5 of the working electrode 1, this layer 6 can be applied in the form of a hardened paste of carbon particles and a polymeric binder. The field of layer 6 preferably has the same size as the field of enzyme layer 5, but this is not essential. However, no measures may be taken to increase the surface of the counter electrode, and the counter electrode 2 may be coated without any kind of coating or optionally made from the block copolymer described above and optionally Together with the spacer, it may be designed to be a linear conductor path.

  The reference electrode 3 is 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 consists of a conductor 3a on which a field 3a of conductive silver / silver chloride paste is arranged.

  FIG. 4 shows a cross-sectional view along the cross-sectional line CC of FIG. The section line CC extends through one of the enzyme layer fields 5 of the working electrode 1 and to between the conductive layer fields 6 of the counter electrode 2. Between the fields of the enzyme layer 5, the conductor 1a of the working electrode 1 may be catalyzed by the metal of the conductive paths 1a, 2a, as well as the conductor 2a of the counter electrode 2 between the fields of the conductive layer 6. In order to prevent reaction, it can be covered by an electrically insulating layer 7. The field of the enzyme layer 5 is therefore located in the opening of the insulating layer 7. Similarly, the field of the conductive layer 6 of the counter electrode 2 can also be located over the opening of the insulating layer 7.

  The enzyme layer 5 is covered by a cover layer 8 which exhibits a diffusion resistance to the analyte to be measured and thus acts as a diffusion barrier. The diffusion barrier 8 consists of a single copolymer having alternating hydrophilic and hydrophobic blocks, as described above.

  A preferred 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. Due to its diffusion resistance, the cover layer 8 allows only a few analyte molecules to reach the enzyme layer 5 per unit time. Thus, the cover layer 8 reduces the rate at which analyte molecules are converted, and thus counters depletion of analyte concentration around the working electrode.

  The cover layer 8 extends continuously over substantially the entire area of the conductor 1a of the working electrode 1. On the cover layer 8, a biocompatible membrane can be placed as a spacer 9 that establishes a minimum distance between the enzyme layer 5 and the cells around the body tissue. This measure is advantageously used for analyte molecules such that in the event of a transient failure of fluid exchange around the enzyme layer field 5, the analyte molecules can reach the corresponding enzyme layer field 5 therefrom. Create a storage tank. If the exchange of bodily fluids around the electrode system is temporarily restricted or prevented, analyte molecules stored in the spacer 9 diffuse into the enzyme layer 5 of the working electrode 1 where the molecules are converted. to continue. Thus, the spacer 9 causes a significant decrease in the analyte concentration, but the corresponding measurement result distortion only occurs after a significantly longer period. In the exemplary embodiment shown, the membrane forming the spacer 9 also covers the counter electrode 2 and the reference electrode 3.

  The spacer membrane 9 may be a dialysis membrane, for example. In this sense, a dialysis membrane is understood to be a membrane through which molecules larger than the maximum size cannot permeate. The dialysis membrane can be manufactured earlier in a separate manufacturing process and then applied during manufacture of the electrode system. The maximum size of molecules that can be permeated by the dialysis membrane is selected such that analyte molecules can pass but larger molecules are retained.

  Alternatively, instead of a dialysis membrane, a coating made from a polymer that is very permeable to the analyte and water, for example based on polyurethane or acrylate, is applied as a spacer membrane 9 on the electrode system. obtain.

  Preferably, the spacer is made of a copolymer of (meth) acrylate. Preferably, the spacer membrane is a copolymer from at least two or three (meth) acrylates. More preferably, the spacer membrane is greater than 50 mol-%, at least 60 mol-%, or at least 70 mol-% hydrophilic monomer units, such as HEMA and / or 2-HPMA, and 40 mol-%. Up to or up to 30 mol-% hydrophobic units such as BUMA and / or MMA. The spacer can be a random or block copolymer. Particularly preferred spacer membranes comprise MMA or BUMA as the hydrophobic component and 2-HEMA and / or 2-HPMA as the hydrophilic component. Preferably, the amount of hydrophilic monomers HEMA and / or HPMA is between 80 mol-% and 85 mol-%, and the amount of hydrophobic component MMA and / or BUMA is between 15 mol-% and 20 mol-%. Between.

  A highly preferred spacer membrane of the present invention is made from Adapt® (BioInteractions Ltd, Reading, England) copolymer. Adapt® includes BUMA as the hydrophobic moiety and 2-HEMA and 2-HPMA as the hydrophilic moiety, where the amount of 2-HEMA hydrophilic monomer is about 80 mol-%.

  The spacer membrane is very permeable to the analyte, i.e. it has a significant sensitivity per unit area of the working electrode, e.g. less than 20% with a layer thickness of less than about 20 μm, preferably less than about 5 μm, Or it reduces to 5% or less. A particularly preferred spacer membrane thickness is about 1 to about 3 μm.

  The enzyme layer 5 of the electrode system can comprise metal oxide particles, preferably manganese dioxide particles, as catalytic redox mediators. Manganese dioxide catalytically converts hydrogen peroxide, such as formed by enzymatic oxidation of glucose and other biological analytes. During the decomposition of hydrogen peroxide, the manganese dioxide particles move electrons to the conductive component of the working electrode 1, such as graphite particles in the enzyme layer 5. The catalytic decomposition of hydrogen peroxide offsets any decrease in oxygen concentration in the enzyme layer 5. Advantageously, this allows the conversion of the analyte to be detected in the enzyme layer 5 not to be limited by the local oxygen concentration. The use of catalytic redox mediators thus counters the distortion of measurement results caused by lower oxygen concentrations. Another advantage of catalytic redox mediators is to prevent the production of hydrogen peroxide at concentrations that can cause cell damage.

  The preferred spacer membrane polymers described herein may be used as an outer coating for the diffusion barrier of the present invention, but in general electrode systems, in particular, electrodes with immobilized enzyme molecules. And an outer coating of the electrode system for measuring the concentration of the analyte under in vivo conditions comprising a diffusion barrier that controls the diffusion of the analyte from the outside of the electrode system into the enzyme molecules. Thus, although the spacer membrane can be placed on the diffusion barrier, the spacer membrane may also be placed directly on the enzyme layer. In this last sense, the spacer membrane can also act as a diffusion barrier itself and reduce the diffusion of analyte molecules into the enzyme layer.

  When the electrode system of the present invention is inserted or implanted into the body, the spacer membrane becomes the boundary between the implanted sensor and the surrounding body fluid or tissue. Consequently, when exposed to body fluids or tissues, the spacer membrane of the present invention must be mechanically strong so that it does not deform or push the sensor. To achieve this goal, regardless of the inherent hydrophilicity of the copolymer, the moisture uptake of the spacer membrane copolymer and the concomitant swelling of the copolymer must be limited.

（式中、ｍ₁およびｍ₂は、それぞれ、乾燥コポリマーおよび上記の測定条件にしたがった水和後のコポリマーの重量を示す）
にしたがい決定される。 Preferably, the relative moisture uptake of the spacer membrane copolymer should not exceed 50 wt%, preferably 40 wt%, more preferably 30 wt%, relative to the total proportion of the copolymer. In the sense of the present invention, the relative water uptake was measured by subjecting the dried copolymer to an excess of phosphate buffer (pH 7.4) at a temperature of 37 ° C. for 48 hours. The relative water uptake (WU%) is preferably the following formula:

(Where m ₁ and m ₂ represent the weight of the dry copolymer and the copolymer after hydration according to the above measurement conditions, respectively)
To be decided.

  The inventors of the present invention have found that a preferred spacer membrane made with the copolymer Adapt® has a pH of 33 ± 1.8% phosphate buffer (pH 7) based on its own weight at 37 ° C. for 48 hours. .4) found to capture. Under the same conditions, the membrane of polymer Lipidure® CM5206 (NOF Corporation, Japan) takes up 157 ± 9.7 wt% phosphate buffer relative to its own weight. The lower moisture uptake of the polymer advantageously improves the mechanical stability of the spacer membrane of the present invention. In contrast, Lipidure® CM5206 exhibits higher moisture uptake and is more fragile, easily deformed or driven, especially when applied over an enzymatic in vivo sensor electrode system. Swells into a gel.

  Further, during insertion and implantation of the electrode system, the spacer is in direct contact with tissue and / or bodily fluids including biomolecules such as proteins and cells, such as interstitial fluid or blood. Advantageously, the spacer membrane should protect the sensor inserted and implanted in the tissue and / or body fluid environment and therefore the body tissue response to the implant should be minimized. In fact, the body's response to the implanted material is known as the “foreign body response” (FBR). With FBR, the body tries to break the implant or, if that is not possible, attempts to create a capsule to separate it from the surrounding tissue (foreign granulomas). The first step in the FBR reaction is the binding of a protein (eg, fibrinogen, albumin, immunoglobulin, complement, etc.) to the previous material, ie the surface of the implant. This protein coat presents binding sites for receptors on immune cells. For example, fibrinogen contains a structural motif that binds to the monosite receptor MAC-1. When fibrinogen binds to the surface of the implant, its conformation changes and the binding site for MAC-1 is exposed. As a result, immune cells such as monocytes are recruited to the implant and activated, secreting enzymes and radicals that attack the implant. In addition, immune cells secrete soluble factors, ie cytokines, in order to recruit and activate other immune cells and thereby amplify the immune response. If the implant cannot be removed, a fibrous capsule is formed by connective tissue cells and proteins. However, this capsule is a diffusion barrier that prevents the analyte from reaching the sensor. In summary, the event of foreign body reaction as described above is thought to interfere with the function of the electrode system and its lifetime in vivo.

  Thus, the improved spacer membrane of the enzymatic in vivo sensor electrode system further provides a reduction in tissue response to the implant and prevents the formation of capsules that separate the sensor from the surrounding tissue and body fluids.

  Accordingly, a further object of the present invention comprises an electrode on which enzyme molecules are immobilized, and preferably a diffusion barrier that controls the diffusion of analytes from body fluids outside the electrode system to the enzyme molecules, wherein the spacer membrane is Electrode system for measuring the concentration of an analyte under in vivo conditions characterized in that it forms at least part of the outer layer of the electrode system, wherein the spacer membrane is a hydrophilic copolymer of acrylic and / or methacrylic monomers And an electrode system in which the polymer contains more than 50 mol% of hydrophilic monomers.

  As mentioned above, the spacer membranes of the present invention are limited to protect the sensor electrode system from protein adsorption that may trigger immune cell responses and may limit or interfere with its in vivo performance. It has the ability to bind proteins. Examples 5 and 6 show that preferred spacer membranes of the invention provide little binding to fibrinogen and prevent conformational changes of fibrinogen that would lead to exposure of the MAC-1 binding motif to monocytes. Is shown. Advantageously, the spacer membrane copolymer material does not activate immune cells by itself. Example 6 can show that the spacer membrane copolymer of the present invention can reduce immune cell activation by the embedded sensor. Furthermore, the spacer membrane is advantageously a biocompatible material, in particular compatible with body fluids such as blood. Example 7 shows that the spacer membrane copolymer of the present invention can prevent hemolysis and complement activation by the embedded sensor. Thus, the spacer membrane of the present invention not only exhibits high mechanical stability, but also has ideal biocompatibility characteristics, which is surprising considering the low water uptake when wet. It is.

  The properties of this embodiment with particular reference to the structure of the electrode system, the analyte and the enzyme molecule are as described herein. The diffusion barrier is preferably as described herein, however, it may also have a different composition or may not be present. In certain preferred embodiments, the diffusion barrier preferably comprises a block copolymer having at least one hydrophilic block and at least one hydrophobic block as described herein.

  In a further preferred embodiment, the diffusion barrier comprises a hydrophilic polyurethane. Hydrophilic polyurethanes used as diffusion membranes can be made by polyaddition of (poly) diisocyanates, preferably 4,4'-methylene-bis- (cyclohexyl isocyanate), to polyalcohols, preferably diol mixtures.

  The components of the diol mixture are preferably polyalkylene glycols such as, for example, polyethylene glycol (PEG) and polypropylene glycol (PPG), and aliphatic diols such as, for example, 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 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 17.5 mol% PPG. In order to increase the hydrophilicity of the polyurethane, the proportion of PEG may be increased, for example, to 4.5-5.5 mol%, preferably to 5 mol% PEG, in order to obtain a very hydrophilic polyurethane. . Various hydrophilic variants of polyurethane can also be mixed to optimize the properties of the diffusion barrier.

  Preferred acrylic and methacrylic monomers for the spacer membrane copolymer are as described herein.

The hydrophilic monomer unit is preferably a hydrophilic (meth) acrylic acid ester, ie an ester having a polar group, ie an ester having an OH, OCH ₃ or OC ₂ H ₅ group in the alcohol part of the ester, a hydrophilic (Meth) acrylamide having an amide (NH ₂ ) or N-alkyl or N, N-dialkylamide group, the alkyl group having 1 to 3 carbon atoms and optionally eg OH, OCH ₃ or OC Suitable with (meth) acrylamide containing hydrophilic groups such as ₂ H ₅ groups and charged groups such as acrylic acid (acrylate) or methacrylic acid (methacrylate) such as anionic or cationic groups Selected from (meth) acrylic acid units. In addition, combinations of monomer units can be used.

Specific examples of preferred monomer units for the hydrophilic block are selected from:
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,
N-alkyl or N, N-dialkylacrylamide and N-alkyl or N, N-dialkylmethylamide, wherein the alkyl group contains 1 to 3 carbon atoms such as methyl, ethyl or propyl,
Acrylic acid (acrylate),
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 unit is preferably selected from hydrophobic acrylic acid and / or methacrylic acid units, or combinations thereof. Preferably, the hydrophobic monomer unit is selected from hydrophobic (meth) acrylic acid esters, such as, for example, esters having an alcohol moiety with 1 to 3 carbon atoms that do not have a hydrophilic group. Specific examples of monomer units for the hydrophobic block are selected from:
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 a preferred embodiment, the hydrophobic block comprises methyl methacrylate (MMA) and n-butyl methacrylate (BUMA).

  The outer spacer membrane covers at least the part of the working electrode that contains the enzyme molecules, and optionally also other parts, such as the counter electrode. If present, the spacer membrane also covers the reference electrode. The spacer membrane preferably covers the entire embedded surface of the electrode system. The spacer membrane preferably covers the working electrode, optionally the counter electrode and, if present, the reference electrode in the form of a continuous layer.

  The electrode system comprising the improved spacer membrane of the present invention may be made part of the sensor, for example by being coupled to a potentiostat and an amplifier for amplification of the measurement signal of the electrode system. The sensor is preferably an enzymatic, non-fluidic (ENF) sensor, more preferably an electrochemical ENF sensor. The electrodes of the electrode system may be disposed on a substrate carrying a potentiostat or may be bonded to a circuit board carrying a potentiostat. Preferably, the sensor is for the measurement of glucose.

  A further subject matter of the present invention relates to the use of hydrophilic copolymers of acrylic and / or methacrylic monomers, wherein the hydrophilic copolymer comprises more than 50 mol% of hydrophilic monomers as a spacer membrane for the enzymatic electrode. The hydrophilic copolymer is preferably as described above. Preferably, the spacer membrane is used to minimize foreign body reaction (FRB) to the enzymatic electrode when it is inserted or implanted in the body.

Example 1 Permeability of an enzymatic, non-fluidic (ENF) sensor with dispersed electrodes for percutaneous implantation with a diffusion layer consisting of one single block copolymer The sensor is 250 μm thick On a prefabricated palladium strip conductor structure on a polyester substrate with The working electrode (WE) and the counter electrode (CE) were placed in a distributed manner (as shown in FIGS. 1-2).

  The CE field was overprinted with carbon paste and the remaining strip conductors were insulated. The field of WE was overprinted with a mixture of cross-linked glucose oxidase (enzyme), conductive polymer paste and electrocatalyst, here manganese dioxide (Technipur). The remaining path of the strip conductor was similarly insulated. The reference electrode (RE) is made of an Ag / AgCl paste. The electrode covers about 1 cm of the sensor shaft.

  The field of WE was coated with a diffusion layer of block copolymer consisting of HEMA block and BUMA block. The layer thickness is 7 μm.

  Four batches of sensors were made, each with a specific block copolymer as a diffusion layer (see list below). All block copolymers are obtained from Polymer Source, Montreal and are listed in Table 1 below.

  Each block copolymer was dissolved in an organic solvent (25% concentration) and used to coat the sensor. After drying using a belt dryer (2 minutes, 30-50 ° C.), the coated sensor was tested in vitro in various concentrations of glucose solution. Ten sensors were measured as random samples for each sensor batch. As a measure of sensitivity in vitro, the signal was calculated by the difference in measured current at 10 mM and 0 mM glucose concentrations and divided by 10 mM (see Example 4).

  All sensors were operated with a polarization voltage of 350 mV for Ag / AgCl and the measurement temperature was kept constant at 37 ° C. The sensor used in this measurement series did not have the spacer described in WO 2010/028708, but did not make any difference in light of the signal level tested. FIG. 5 shows sensor sensitivity using standard deviation for four different diffusion layers.

  For block copolymers C, D and F, there is a clear link between in vitro sensitivity and the molar ratio of hydrophobic block to hydrophilic block. In almost the same chain length copolymer, the sensitivity improved as the amount of hydrophilic block (HEMA) increased.

  Sensors with a diffusion layer of block copolymer B are an exception. Although polymer B has a relative ratio of hydrophobic amount to hydrophilic amount, similar to polymer F, sensitivity and thus glucose permeability is reduced. Empirically, it can be said that in the case of polymer B, the total chain length (corresponding to the molecular weight (total) of the copolymer molecules) is very large, which reduces the permeability of the layer. This may also be seen in the moisture uptake of block copolymer B compared to other polymers, measured gravimetrically. Polymer B has a moisture uptake (weight percent relative to the dry weight of the polymer) of 10.6% + 1.5%. Polymer C is 15.6% + 0.0%, Polymer F is 16.5% + 3.1%, and Polymer D is 27% + 1.7%.

Example 2 Mechanical Flexibility of ENF Glucose Sensor Diffusion Layer A sensor was made with the diffusion layer of the present invention, however, as described in WO 2010/028708. It was assumed that the glass transition temperature (Tg) is an alternative parameter of mechanical flexibility. Furthermore, it was hypothesized that the glass transition temperature, which can be assigned to the hydrophobic block, determines the mechanical flexibility in in vivo applications. It should be noted that for one block copolymer, several Tg's may be specified corresponding to the number of blocks.

  The sensor was coated using the same electrode paste as in Example 1. Several sensors were then coated with a copolymer selected from MMA-HEMA (manufactured by Polymer Source, Montreal). This polymer (referred to as E) has a total molecular weight of 41 kD and the molar ratio of MMA (hydrophobic amount) to HEMA is 60%: 40%. The glass transition temperature of the hydrophobic block, measured at a heating rate of 10 ° C./min using DSC, is 111 ° C.

  The other sensor was provided with a diffusion layer of the block copolymer of the present invention (referred to as A). The hydrophobic block of copolymer A comprises MMA and BUMA in equimolar amounts within a random sequence. Again, the molar ratio of the hydrophobic portion to the hydrophilic portion is 60%: 40%. The molecular weight is 36 kD. The Tg of the hydrophobic block drops to 73 ° C. due to the random sequence of MMA and BUMA (about 45 ° C. Tg).

  Both diffusion layers were made from the respective ether solution (25%) of the copolymer and dried as described in Example 1. The thickness of the diffusion layer was 7 μm. The spacer layer was subsequently applied by dip coating and dried at room temperature for 24 hours. The spacer layer was made from Lipidure® CM5206 manufactured by NOF Japan.

  After explanting from the tissue, the sensor with the copolymer E diffusion layer shows sporadic cracks in the diffusion layer. This is considered as a mechanical load effect. In contrast, a sensor with a copolymer A diffusion layer does not exhibit any cracks under the same load. This is apparently due to a decrease in Tg, which improves the mechanical stability of the copolymer. Physical mixing of the two copolymers, as disclosed in WO2010 / 028708, is no longer necessary.

Example 3 Optimized permeation behavior of an ENF glucose sensor with dispersed electrodes and diffusion layers of the present invention The sensor is as described in Example 1, but with an additional spacer layer over the entire sensor shaft It was manufactured to be equipped with. Sensors were produced having respective diffusion layers that were copolymers A, C, D and F as described in Examples 1 and 2. For this purpose, a 24% ether solution of the copolymer was produced. Each solution was applied on a set (N = 10) of sensors and dried in a band dryer. Thereby, a diffusion layer having a thickness of 7 μm was obtained.

  The sensor was then provided with a spacer layer as described in Example 2.

  The sensor is coupled to the measurement system on the upper end of the sensor, which moves the measured data to the data store. In vitro measurements were performed as in Example 1, however, with a measurement period of 7 days. From the measured data, sensitivity drift was calculated over the corresponding measurement period for each sensor. FIG. 6 shows the average value of in vitro drift values for each group for each variant of the sensor, i.e. a sensor with a different diffusion layer. The initial stage of measurement (the first 6 hours, the so-called measurement start stage) was excluded from the calculation.

  For all copolymers C, D and F with a hydrophobic block of BUMA, there is a positive drift, ie the sensitivity improves with time. In contrast, copolymer A, which has a hydrophobic block of random copolymers of MMA and BUMA, results in a very low, slightly negative drift.

  These differences may be explained by the long-term transmission responsiveness of each diffusion layer, measured by additional experiments. A palladium sensor within the WE paste but having a defined activated surface, ie without an enzyme layer (excluding the effect of its expansion behavior on the result) is coated with the polymer solution described above, Then, it was dried and the layer thickness was measured. Subsequently, the conductivity was measured in a sodium and chlorine containing buffer.

  FIG. 7 shows that the conductivity of copolymer A was kept approximately constant after the short measurement start phase.

  In the case of copolymer F, this was not the case as shown in FIG. 8 even under identical measurement conditions. In this case, a long-term and strong transmission response of the copolymer F diffusion layer was observed, which was substantially unaffected by the layer thickness. In the case of copolymer F (and in the case of copolymers C and D (data not shown)) with a BUMA hydrophobic block, an increase in permeability is provided even over a long period of time. If measured, if a diffusion layer was applied on a sensor with a dispersed enzyme layer, this leads to a continuous increase in sensitivity. This explains the observed positive sensor drift.

  Similarly, the sensor with block copolymer A exhibits a negative drift, which is due to a very small permeability change in the conductivity measurement. However, a large increase in conductivity is observed in copolymer A immediately after the start of the measurement (up to about 1 hour thereafter). Here, a very fast rise is observed, which ends after about 1 hour. At this point, the diffusion layer is completely wetted and the structural reconstruction due to moisture uptake ends. The degree of structural change is probably dependent on Tg. It seems reasonable that a copolymer with increased Tg passes through a reconstruction that is limited in time and size compared to a copolymer with a Tg in the range of atmospheric temperature.

  In addition, it should be mentioned that the sensor with copolymer A exhibits a relatively high sensitivity at the start of the measurement compared to a sensor with a diffusion layer of copolymer F. This is expected due to the same relative ratio between the hydrophobic and hydrophilic blocks. The resulting sensitivity range of 1-1.5 nA / mM (see Example 1) appears to be ideal. This sensitivity is likewise obtained in sensors having a diffusion layer made of copolymer A.

  With respect to the sum of the three physicochemical properties (permeability, mechanical stability and permeation responsiveness), the ideal sensor is preferably hydrophobic with at least two different hydrophobic monomer units arranged in a random fashion It can be obtained using a diffusion layer such as a block copolymer with blocks, for example block copolymer A. None of the other block copolymers whose hydrophobic blocks consist only of a single monomer unit reach a quality comparable to copolymer A in all three parameters.

Example 4 Characterization of Block Copolymers Multiple field sensors (10 fields each for working and counter electrodes) for continuous glucose measurement were made and characterized in vitro.

  The sensor has a diffusion layer consisting of a block copolymer comprising a hydrophobic block of methyl methacrylate (MMA) and n-butyl methacrylate (BUMA) randomly copolymerized and a hydrophilic block of 2-hydroxyethyl methacrylate (HEMA). Provided. These polymers (defined as G and H) are manufactured by Polymer Source, Montreal and are more permeable than polymer A from Examples 1-3, which is hereby incorporated by reference. Embedded in.

  In Table 2 below, the copolymers are described.

  The molecular weight Mn of each block is shown individually in Table 2 above, and since the polymer is known to have a molecular chain length dispersion around a certain average value, the average value is shown. Yes. This also applies for the quantities derived in Table 2.

  The indicated glass transition temperature of the hydrophobic block is within the preferred range to ensure mechanical flexibility.

個々の測定値（Ｎ＝８）から平均感度ＳＥ_mが決定された。得られた感度値は、顕微鏡で測定された、複数のフィールドを有するセンサー上の全ての作用電極スポットの幾何学的な総面積Ｆで除された。これにより、感度密度ＳＥ_m／Ｆが得られた。 An important parameter regarding the permeability of the diffusion barrier for the analyte is the sensitivity per unit area (ie, geometric area) of the working electrode. The sensitivity SE was calculated for each sensor analyzed at nA / mM by current (I) measurements at 10 mM and 0 mM glucose concentrations in phosphate buffer (pH 7.4):

The average sensitivity SE _m was determined from the individual measurements (N = 8). The sensitivity value obtained was divided by the geometric total area F of all working electrode spots on the sensor with multiple fields, measured with a microscope. Thereby, a sensitivity density SE _m / F was obtained.

にしたがい、分析されたセンサーのそれぞれについて算出された。個々の測定値から、平均線形性値およびその標準偏差が決定された（表３参照のこと）。 The linearity Y of the in vitro function curve indicates the diffusion control functionality of the polymer cover layer on the working electrode. From amperometric measurements at 20 mM, 10 mM and 0 mM glucose concentrations, the following formula:

Therefore, it was calculated for each of the analyzed sensors. From the individual measurements, the mean linearity value and its standard deviation were determined (see Table 3).

にもとづき、ｃｍ²／ｓで算出され得、ここで式中、ＳＥ_mおよびＬ_mは感度および層厚それぞれの平均値であり、および、Ｆは全ての作用電極スポットの総面積である。 Finally, the layer thickness L of the sensor diffusion barrier was measured by optical measurement of the respective polymer. Corresponding average values were calculated for ≧ 23 sensor samples with the same polymer. From these, the effective diffusion coefficient D _eff of the cover layer is given by the following formula:

Based on the above, it can be calculated in cm ² / s, where SE _m and L _m are average values of sensitivity and layer thickness, respectively, and F is the total area of all working electrode spots.

  Sensor drift was calculated from repeated glucose concentration steps of in vitro measurements over 7 days. The results for polymer H, which shows a substantially constant conductivity, are shown in FIG.

  Table 3 below shows the results of functional characterization.

For the more hydrophilic polymer G (which is more permeable to glucose), the diffusion coefficient is also an alternative method, for example, the polymer from a chamber containing a volcose solution to a chamber containing a glucose-free buffer. Measured using glucose permeation through the film. By this method, a similar value was obtained as the diffusion coefficient (1.17 × 10 ⁻⁹ cm ² / s).

Example 5 Binding of Protein to Spacer Layer Material To evaluate protein binding to spacer layer material, Adapt® (BioInteractions Ltd, Reading, England) or Eudragit® E100 (Evonik Industries) Filled into incubation plates (FluoroNunc Maxisorp, Thermo Scientific). Eudragit® E100 is a cationic copolymer based on dimethylaminoethyl methacrylate, butyl methacrylate and methyl methacrylate. The polymer was dried overnight at 40 ° C. The spacer material was then covered with a fibrinogen solution. The solution contained fibrinogen derived from human plasma and bound to the fluorescent dye Alexa488 (purchased from Invitrogen). After 4 hours of incubation, the fibrinogen solution was aspirated and the spacer layer was washed 8 times with borate buffer. The amount of protein bound to the spacer was analyzed by measuring the fluorescence intensity in the incubation plate using a fluorescence reader (Synergy 4, BioTek Instruments) at an excitation wavelength of 485 nm and an emission wavelength of 528 nm. A known concentration of labeled protein (6.25-500 ng) was used to generate a calibration curve to convert fluorescence readings into protein quantities.

  As expected, fibrinogen adhered to the uncoated incubation plate (blank), resulting in 390 ng binding protein (FIG. 10). Plates coated with Eudragit® E100 showed reduced protein binding, which was 60 ng. Little protein binding was detected in the Adapt®-coated plates. The pre-incubation reading was due to background fluorescence. These results clearly show that the surface coated with the spacer material, in particular the surface coated with Adapt®, well prevents fibrinogen adhesion.

Example 6 Cytokine release from cells after contact with a spacer layer A sensor was prepared as described in Example 2. The sensor was then provided with a spacer layer as described in Example 2. The spacer layer was made from Lipidure® CM5206 (NOF Corporation, Japan) or from Adapt® (BioInteractions Ltd, Reading, England).

  A sensor without a spacer layer, a sensor with a spacer layer consisting of Lipidure® CM5206 or a sensor with a spacer layer consisting of Adapt® is incubated with monocytic THP-1 cells and inflammation Sex marker induction was analyzed.

  THP-1 cells were cultured for 24 hours at 37 ° C. in the presence of the sensor. The cells were then collected by centrifugation. The supernatant is used to measure cytokine release, while the cell pellet is resuspended in PBS containing 1% bovine serum albumin (BSA) and is also known as cell surface protein CD54 (ICAM-1) Used to analyze the expression of inflammatory biomarkers). THP-1 cells were incubated with an anti-CD54 antibody (BD Bioscience) labeled with the fluorescent dye phycoerythrin. After 45 minutes incubation at 4 ° C., the cells were washed in PBS / 1% BSA and the mean fluorescence intensity (MFI) of 10000 cells was determined using a flow cytometer (BD FACSArray, BD Bioscience). Measured (excitation wavelength 532 nm, emission wavelength 585 nm). Compared to untreated THP-1 cells, incubation with the uncoated sensor increased CD54 relative expression (6-fold induction) as shown by high MFI readings (FIG. 11). Brought about. Incubation of cells with sensors coated with CM5256 or Adapt® spacer layer resulted in a 45% or 41% decrease in CD54 expression, respectively.

  The supernatant should be in accordance with the manufacturer's instructions for the amount of cytokines, interleukin-8 (IL-8) and “monocyte chemotactic protein-1” (MCP-1). It was used to measure using immunoassay with beads (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).

  Compared to untreated THP-1 cells, the uncoated sensor produced potent IL-8 (49 vs 197 pg / mL) (Figure 12a) and MCP-1 (6 vs 48 pg / mL) (Figure 12b). Release was induced. Release of IL-8 and MCP-1 was reduced when the sensor was covered with CM5256 or Adapt® spacer layer. The sensor coated with Adapt® induced the release of 100 pg / mL IL-8 and 25 pg / mL MCP-1. A sensor coated with CM5256 resulted in secretion of 125 pg / mL IL-8 and 18 pg / mL MCP-1.

  Taken together, these data indicate that the spacer layer reduces the induction of three well-known inflammatory biomarkers, CD54, IL-8 and MCP-1.

  In order to analyze the role of protein adsorption for activation of THP-1 cells, tissue culture plates were coated with CM5256, Adapt® or Eudragit® E100. The spacer was then incubated with human fibrinogen (Sigma-Aldrich). THP-1 cells were incubated with various spacers + fibrinogen layers. After 48 hours of incubation at 37 ° C., the cells were precipitated by centrifugation and the supernatant was analyzed for IL-8 release. As a control, cells were cultured in culture plates (Polyst. = Culture plate material) without spacers but coated with fibrinogen. As shown in FIG. 13, these 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 cultured on Eudragit® E100 + fibrinogen released 206 pg / mL IL-8. In particular, cells cultured 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 probably exposes the MAC-1 binding site. THP-1 cells activated via binding to the MAC-1 receptor release cytokines such as IL-8 and thereby trigger an inflammatory response. Therefore, a spacer layer made of Adapt® or CM5256 avoids protein deposition and exposure of structural motifs on the surface (such as sensors) and thereby reduces the inflammatory response to the implant. Minimize.

Example 7 Limited Hemolysis with Sensor Coated with Spacer Layer A sensor was prepared as described in Example 2. The sensor was then provided with a spacer layer as described in Example 2. The spacer layer was made from Lipidure® CM5206 (NOF Corporation, Japan) or from Adapt® (BioInteractions Ltd, Reading, England).

The possibility of hemolysis of a sensor without a spacer layer, a sensor with a spacer layer consisting of Lipidure® CM5206 or a sensor with a spacer layer consisting of Adapt® was analyzed. For this purpose, a sensor with a total surface area of 6 cm ² was incubated with red blood cells and then the lysis of the cells was measured by measuring the release of hemoglobin into the supernatant. Red blood cells were isolated from fresh human blood by centrifugation (citrate was used to avoid aggregation). They were then washed with phosphate buffered saline (PBS) and then diluted 1:40 in PBS. The erythrocyte suspension was incubated with the sensor for 24 hours on a rotating platform (350 rpm) at 37 ° C. in the dark. The cells were then sedimented by centrifugation and the hemoglobin content of the supernatant was determined spectrally by measuring the absorbance of the supernatant at a wavelength of 575 nm. The results are indicated by the solubility index in%, which is the amount of hemoglobin released in the sample divided by the amount of hemoglobin released in the positive control (= complete osmotic lysis of red blood cells in distilled water). It is. The result is shown in FIG.

  The sensor without the spacer layer caused significant hemolysis as indicated by a high hemolysis index of 47.4%. Coating the sensor with Lipidure® CM5206 spacer layer reduced the sensor's hemolytic potential as indicated by a solubility index of 14.7%. A sensor coated with an Adapt® spacer layer caused a slight hemolysis resulting in a solubility index of 7.5%, which was a negative control (= in PBS without any test material) Red blood cells) or Adapt® only. These results indicated the protective function of the spacer layer to reduce hemolysis.

Claims (28)

An electrode system for measuring the concentration of an analyte under in vivo conditions comprising an electrode with immobilized enzyme molecules comprising:
A spacer membrane forms at least a portion of the outer layer of the electrode system, the spacer membrane is formed from a hydrophilic copolymer of acrylic and / or methacrylic monomers, and the hydrophilic copolymer is 50 mol-%. An electrode system comprising more hydrophilic monomer units . The electrode system according to claim 1, further comprising a diffusion barrier that controls diffusion of the analyte from outside the electrode system to the enzyme molecules. The electrode system according to claim 1 or 2, wherein the spacer membrane is a hydrophilic copolymer from at least two or three acrylic and / or methacrylic monomers. The hydrophilic monomer unit is selected from hydrophilic (meth) acrylic acid ester having a polar group, hydrophilic (meth) acrylamide, (meth) acrylic acid, or a combination thereof. 2. The electrode system according to item 1. The electrode system according to claim 4, wherein the polar group is selected from OH, OCH ₃ or OC ₂ H ₅ groups. The hydrophilic monomer unit is
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,
N-alkyl or N, N-dialkylacrylamide and N-alkyl or N, N-dialkylmethylamide,
The alkyl group having 1 to 3 carbon atoms,
Acrylic acid,
6. An electrode system according to claim 4 or 5 selected from methacrylic acid and combinations thereof. The electrode system according to claim 6, wherein the hydrophilic monomer unit is selected from 2-hydroxyethyl methacrylate (HEMA), 2-hydroxypropyl methacrylate (2-HPMA), and combinations thereof. The electrode system according to claim 1, wherein the hydrophilic copolymer of the spacer membrane comprises at least 60 mol-% hydrophilic monomer units . The electrode system according to claim 1, wherein the hydrophilic copolymer of the spacer membrane comprises at least 70 mol-% of hydrophilic monomer units . The electrode system according to any one of claims 1 to 9, wherein the hydrophilic copolymer of the spacer membrane further comprises 40 mol-% or less of hydrophobic monomer units . The electrode system according to any one of claims 1 to 9, wherein the hydrophilic copolymer of the spacer membrane further comprises 30 mol-% or less of a hydrophobic monomer unit . The hydrophobic monomer unit is
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,
12. An electrode system according to claim 10 or 11 selected from neopentyl methacrylate and combinations thereof. 13. The electrode system according to claim 12, wherein the hydrophobic monomer unit is selected from methyl methacrylate (MMA), n-butyl methacrylate (BUMA) and combinations thereof. The hydrophobic monomer unit is methyl methacrylate (MMA) or n-butyl methacrylate (BUMA), and the hydrophilic monomer unit is 2-hydroxyethyl methacrylate (HEMA) and / or 2-hydroxypropyl methacrylate (2 The electrode system according to claim 10, which is -HPMA). The spacer membrane is an electrode system formed from a copolymer comprising monomer units of n-butyl methacrylate (BUMA), 2-hydroxyethyl methacrylate (2-HEMA) and 2-hydroxypropyl methacrylate (2-HPMA). The electrode system according to claim 10, wherein the copolymer comprises 80 mol-% 2-HEMA monomer units . The electrode system according to claim 1, wherein the hydrophilic polymer is a random copolymer or a block copolymer. The electrode system according to claim 1, wherein the spacer membrane has a thickness of less than 20 μm. The electrode system according to claim 17, wherein the spacer membrane has a thickness of less than 5 μm. The electrode system according to claim 17, wherein the spacer membrane has a thickness of 1 to 3 μm. 20. The electrode system according to any one of claims 1 to 19, wherein the relative water uptake of the hydrophilic copolymer does not exceed 50% by weight relative to the total weight of the copolymer. 21. The electrode system of claim 20, wherein the relative water uptake of the hydrophilic copolymer does not exceed 40% by weight relative to the total weight of the copolymer. 21. The electrode system of claim 20, wherein the relative water uptake of the hydrophilic copolymer does not exceed 30% by weight relative to the total weight of the copolymer. 23. The electrode system according to any one of claims 2 to 22, wherein the diffusion barrier comprises a hydrophilic polyurethane or a block copolymer having at least one hydrophilic block and at least one hydrophobic block. A counter electrode (2) having a conductor (2a) and a working electrode (1) having a conductor (1a) are included on the substrate (4), and further fixed on the working electrode (1). An enzyme layer (5) containing the enzyme molecules, a diffusion barrier (8) covering the enzyme layer (5), and a spacer membrane (9) covering the diffusion barrier (8) and the counter electrode (2) The electrode system according to any one of claims 1 to 23. 25. A sensor that is insertable or implantable into the body, comprising an electrode system according to any one of claims 1-24. 26. A sensor according to claim 25 for measuring glucose. Use of a hydrophilic copolymer of acrylic and / or methacrylic monomers, said hydrophilic copolymer comprising more than 50 mol-% hydrophilic monomer units as a spacer membrane for an enzymatic electrode system. 28. Use according to claim 27 for minimizing foreign body reaction (FRB) to an enzymatic electrode system.

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