DIFFUSION LAYER FOR AN ENZYMATIC IN-VIVO SENSOR

Abstract

The present disclosure relates to an electrode system for measuring the concentration of an analyte under in-vivo conditions, comprising an electrode with immobilized enzyme molecules and a diffusion barrier that controls diffusion of the analyte from body fluid surrounding the electrode system to the enzyme molecules.

Claims

1. An electrode system for measuring the concentration of an analyte under in-vivo conditions, comprising: an electrode having immobilized enzyme molecules and a diffusion barrier configured to control diffusion of the analyte from the exterior of the electrode system to the enzyme molecules; and the diffusion barrier comprising a block copolymer having from 75 to 95 mol % of at least one hydrophilic block and from 5 to 25 mol % of at least one hydrophobic block, based on the total amount (mol) of the block copolymer in a dry state; wherein the glass transition temperature of the at least one hydrophobic block is in the range of about 80 C. and about 0 C.

2. The electrode system of claim 1, wherein the block copolymer is a diblock copolymer or a triblock copolymer.

3. The electrode system of claim 1, wherein: (i) the hydrophilic block of the block copolymer has a chain length selected from 150 to 900, 200 to 800, 175 to 450, and 300 to 600 monomeric units; and/or (ii) the hydrophobic block of the block copolymer has a chain length selected from 20 to 300, 50 to 250, 50 to 100 and 50 to 150 monomeric units.

4. The electrode system of claim 1, wherein the relative ratio of the chain length of the at least one hydrophilic block and the chain length of the at least one hydrophobic block is in a range selected from 2:1 to 20:1, 3:1 to 10:1, and 4:1 to 8:1.

5. The electrode system of claim 1, wherein a hydrophobic block is made from monomeric units selected from hydrophobic (meth)acrylesters, or combinations thereof.

6. The electrode system of claim 5, wherein the monomeric units are selected from the group consisting of: ethyl acrylate (EA), n- or i-propyl acrylate, n/sec or i-butyl acrylate (n-BuA/sec-BuA or i-BuA), 3-pentyl acrylate, n-pentyl acrylate, n-hexyl acryl ate, n-hexyl methacryl ate, 2-ethylhexyl acryl ate, 2-ethylhexyl methacryl ate, n-heptyl acrylate, n-heptyl methacrylate, n-octyl acryl ate, n-octyl methacrylate, n-nonyl acrylate, n-nonyl methacrylate, n/iso-decyl methacrylate, n-dodecyl methacrylate, n-dodecyl acrylate and combinations thereof.

7. The electrode system of claim 1, wherein the at least one hydrophobic block consists of a homopolymer of a single hydrophobic monomeric unit.

8. The electrode system of claim 1, wherein the at least one hydrophobic block has a weight average molecular weight (Mw) of about 19000 g/mol or lower.

9. The electrode system of claim 8, wherein the at least one hydrophobic block has a weight average molecular weight (Mw) in the range of about 5000 g/mol to about 16000 g/mol.

10. The electrode system of claim 1, wherein the at least one hydrophilic block is made from hydrophilic monomeric units selected from hydrophilic (meth)acrylesters with a polar OH and/or OCH3 group, hydrophilic (meth)acrylamides, (meth)acrylic acid or combinations thereof.

11. The electrode system of claim 10, wherein the at least one hydrophilic block is made from hydrophilic monomeric units selected from group consisting of: 2-hydroxyethyl acrylate, 2-hydroxyethyl methacrylate (HEMA), 2-methoxyethyl acrylate, 2-methoxyethyl 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, 1- or 2-glycerol acrylate, 1- or 2-glycerol methacrylate, acrylamide, methacrylamide, an N-alkyl- or N,N-dialkyl acrylamide, and an N-alkyl- or N,N-dialkyl methacrylamide, wherein alkyl comprises 1 C-atom such as methyl, acrylic acid (acrylate), methacrylic acid (methacrylate), 2-(2-methoxyethoxy)ethylmethacrylate, 2-(2-methoxyethoxy)ethylacrylate, 2-(2-hydroxyethoxy)ethylacrylate, 2-(2-hydroxyethoxy)ethylmethacrylate, and combinations thereof.

12. The electrode system of claim 1, wherein the at least one hydrophilic block has a weight average molecular weight (Mw) of about 24000 g/mol or higher.

13. The electrode system of claim 12, wherein the at least one hydrophilic block has a weight average molecular weight (Mw) in the range of about 24000 g/mol to about 90000 g/mol.

14. The electrode system of claim 1, wherein the hydrophilic block has a glass transition temperature selected from the group consisting of: about 100 C. or above, from 105 C. to 140 C., and from 110 C. to 120 C.

15. The electrode system of claim 1, wherein the molar ratio of hydrophilic block: hydrophobic block is in the range of >75 mol-% (hydrophilic), <25 mol-% (hydrophobic) to 95 mol-% (hydrophilic), 5 mol-% (hydrophobic).

16. The electrode system of a claim 1, wherein the block copolymer has a number average molecular weight (Mn) in the range from 20 to 200 kDa.

17. The electrode system of a claim 16, wherein the block copolymer has a number average molecular weight (Mn) in the range from 25 to 95 kDa.

18. A sensor configured for insertion or implantation into a body, comprising the electrode system of claim 1.

19. The sensor of claim 18, wherein the sensor is configured for the measurement of glucose.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0115] The above-mentioned aspects of exemplary embodiments will become more apparent and will be better understood by reference to the following description of the embodiments taken in conjunction with the accompanying drawings, wherein:

[0116] FIG. 1 shows an exemplary embodiment of an electrode system according to this disclosure;

[0117] FIG. 2 shows a detail view of FIG. 1;

[0118] FIG. 3 shows another detail view of FIG. 1;

[0119] FIG. 4 shows a section along the section line CC of FIG. 2;

[0120] FIG. 5a shows the effective diffusion coefficient (with standard deviation) of three glucose sensors (at 10 mM glucose) provided with different block copolymers (A, B and C) as diffusion membranes;

[0121] FIG. 5b shows the sensor drift of three glucose sensors provided with different block copolymers (A, B and C) as diffusion membranes;

[0122] FIG. 6 shows the results of bending tests of glucose sensors provided with the block copolymer B (Poly(nBuA-b-HEMA)) and a comparative block copolymer Poly(BUMA-r-MMA-b-HEMA);

[0123] FIG. 7a shows the diffusion coefficient (with standard deviation) of a glucose sensor (at 10 mM glucose) provided with a block copolymer E as diffusion membrane;

[0124] FIG. 7b shows the sensor drift of the glucose sensor provided with the block copolymer E as diffusion membrane; and

[0125] FIGS. 8a and 8b show the conductivity of the block copolymers B and E, respectively, dependent on time.

DESCRIPTION

[0126] The embodiments described below are not intended to be exhaustive or to limit the invention to the precise forms disclosed in the following detailed description. Rather, the embodiments are chosen and described so that others skilled in the art may appreciate and understand the principles and practices of this disclosure.

[0127] FIG. 1 shows an exemplary embodiment of an electrode system for insertion into body tissue of a human or animal, for example into cutis or subcutaneous fatty tissue. A magnification of detail view A is shown in FIG. 2, a magnification of detail view B is shown in FIG. 3. FIG. 4 shows a corresponding sectional view along the section line, CC, of FIG. 2.

[0128] The electrode system shown has a working electrode 1, a counter electrode 2, and a reference electrode 3. Electrical conductors of the electrodes 1a, 2a, 3a are arranged in the form of metallic conductor paths, preferably made of palladium or gold, on a substrate 4. In the exemplary embodiment shown, the substrate 4 is a flexible plastic plate, for example made of polyester. The substrate 4 is less than 0.5 mm thick, for example 100 to 300 micrometres (m), and is therefore easy to bend such that it can adapt to movements of surrounding body tissue after its insertion. The substrate 4 has a narrow shaft for insertion into body tissue of a patient and a wide head for connection to an electronic system that is arranged outside the body. The shaft of the substrate 4 preferably is at least 1 cm in length, in particular 2 cm to 5 cm.

[0129] In the exemplary embodiment shown, one part of the measuring facility, namely, the head of the substrate, projects from the body of a patient during use. Alternatively, it is feasible just as well, though, to implant the entire measuring facility and transmit measuring data in a wireless fashion to a receiver that is arranged outside the body. The working electrode 1 carries an enzyme layer 5 that contains immobilized enzyme molecules for catalytic conversion of the analyte. The enzyme layer 5 can be applied, for example, in the form of a curing paste of carbon particles, a polymeric binding agent, a redox mediator or an electro-catalyst, and enzyme molecules. Details of the production of an enzyme layer 5 of this type are disclosed, for example, in WO 2007/147475, reference to which is be made in this context.

[0130] In the exemplary embodiment shown, the enzyme layer 5 is not applied continuously on the conductor 1a of the working electrode 1, but rather in the form of individual fields that are arranged at a distance from each other. The individual fields of the enzyme layer 5 in the exemplary embodiment shown are arranged in a series.

[0131] The conductor 1a of the working electrode 1 has narrow sites between the enzyme layer fields that are seen particularly well in FIG. 2. The conductor 2a of the counter electrode 2 has a contour that follows the course of the conductor 1a of the working electrode 1. This means results in an intercalating or interdigitated arrangement of working electrode 1 and counter electrode 2 with advantageously short current paths and low current density.

[0132] In order to increase its effective surface, the counter electrode 2 can be provided with a porous electrically conductive layer 6 that is situated in the form of individual fields on the conductor 2a of the counter electrode 2. Like the enzyme layer 5 of the working electrode 1, this layer 6 can be applied in the form of a curing paste of carbon particles and a polymeric binding agent. The fields of the layer 6 preferably have the same dimensions as the fields of the enzyme layer 5, although this is not obligatory. However, measures for increasing the surface of the counter electrode can just as well be foregone and the counter electrode 2 can just as well be designed to be a linear conductor path with no coatings of any kind, or with a coating made from the described block copolymer and optionally a spacer.

[0133] The reference electrode 3 is arranged between the conductor 1a of the working electrode 1 and the conductor 2a of the counter electrode 2. The reference electrode shown in FIG. 3 consists of a conductor 3a on which a field 3b of conductive silver/silver chloride paste is arranged.

[0134] FIG. 4 shows a schematic sectional view along the section line, CC, of FIG. 2. The section line, CC, extends through one of the enzyme layer fields 5 of the working electrode 1 and between the fields of the conductive layer 6 of the counter electrode 2. Between the fields of enzyme layer 5, the conductor 1a of the working electrode 1 can be covered with an electrically insulating layer 7, like the conductor 2a of the counter electrode 2 between the fields of the conductive layers 6, in order to prevent interfering reactions which may otherwise be catalysed by the metal of the conductor paths 1a, 2a. The fields of the enzyme layer 5 are therefore situated in openings of the insulation layer 7. Likewise, the fields of the conductive layer 6 of the counter electrode 2 may also be placed on top of openings of the insulation layer 7.

[0135] The enzyme layer 5 is covered by a cover layer 8 which presents a diffusion resistance to the analyte to be measured and therefore acts as a diffusion barrier. The diffusion barrier 8 consists of a single copolymer with alternating hydrophilic and hydrophobic blocks as described above.

[0136] A favorable thickness of the cover layer 8 is, for example, 5 to 60 m, particularly from about 10 to 125 m or from about 10 to 15 m. Because of its diffusion resistance, the cover layer 8 causes fewer analyte molecules to reach the enzyme layer 5 per unit of time. Accordingly, the cover layer 8 reduces the rate at which analyte molecules are converted, and therefore counteracts a depletion of the analyte concentration in surroundings of the working electrode.

[0137] The cover layer 8 extends continuously essentially over the entire area of the conductor 1a of the working electrode 1. On the cover layer 8, a biocompatible membrane may be arranged as spacer 9 that establishes a minimal distance between the enzyme layer 5 and cells of surrounding body tissue. This means advantageously generates a reservoir for analyte molecules from which analyte molecules can get to the corresponding enzyme layer field 5 in case of a transient disturbance of the fluid exchange in the surroundings of an enzyme layer field 5. If the exchange of body fluid in the surroundings of the electrode system is transiently limited or even prevented, the analyte molecules stored in the spacer 9 keep diffusing to the enzyme layer 5 of the working electrode 1 where they are converted. The spacer 9 therefore causes a notable depletion of the analyte concentration and corresponding falsification of the measuring results to occur only after a significantly longer period of time. In the exemplary embodiment shown, the membrane forming the spacer 9 also covers the counter electrode 2 and the reference electrode 3.

[0138] The spacer membrane 9 can, for example, be a dialysis membrane. In this context, a dialysis membrane is understood to be a membrane that is impermeable for molecules larger than a maximal size. The dialysis membrane can be prefabricated in a separate manufacturing process and may then be applied during the fabrication of the electrode system. The maximal size of the molecules for which the dialysis membrane is permeable is selected such that analyte molecules can pass, while larger molecules are retained.

[0139] Alternatively, instead of a dialysis membrane, a coating made of a polymer that is highly permeable for the analyte and water, for example on the basis of polyurethane or of acrylate, can be applied over the electrode system as spacer membrane 9.

[0140] Preferably, the spacer membrane is made from a copolymer of (meth)acrylates. Preferably, the spacer membrane is a copolymer from at least 2 or 3 (meth)acrylates. More preferably, the spacer membrane comprises more than 50 mol-%, at least 60 mol-% or at least 70 mol-% hydrophilic monomer units, e.g., HEMA and/or 2-HPMA, and up to 40 mol-% or up to 30 mol-% hydrophilic units, e.g., BUMA (butylmethacrylate) and/or MMA (methylmethacrylate) based on the total amount of spacer membrane, e.g., as described in WO 2012/130841 and WO 2013/144255. To the person skilled in the art it is clear that the spacer membrane is different from the diffusion barrier. The spacer membrane may be a random or block copolymer. An especially preferred spacer membrane comprises MIVIA or BUMA as hydrophobic moieties and 2-HEMA and/or 2-HPMA as hydrophilic moieties. The spacer membrane is highly permeable for the analyte, i.e., it does significantly lower the sensitivity per area of the working electrode, for example 20% or less, or 5% or less with a layer thickness of less than about 20 m, preferably less than about 5 m. An especially preferred thickness of the spacer membrane is from about 1 m to about 3 m.

[0141] The enzyme layer 5 of the electrode system can contain metal oxide particles, preferably manganese dioxide particles, as catalytic redox mediator. Manganese dioxide catalytically converts hydrogen peroxide that is formed, for example, by enzymatic oxidation of glucose and other bioanalytes. During the degradation of hydrogen peroxide, the manganese dioxide particles transfer electrons to conductive components of the working electrode 1, for example to graphite particles in the enzyme layer 5. The catalytic degradation of hydrogen peroxide counteracts any decrease of the oxygen concentration in the enzyme layer 5. Advantageously this allows the conversion of the analyte to be detected in the enzyme layer 5 to not be limited by the local oxygen concentration. The use of the catalytic redox mediator therefore counteracts a falsification of the measuring result by the oxygen concentration being low. Another advantage of a catalytic redox mediator is that it prevents the generation of cell-damaging concentrations of hydrogen peroxide.

[0142] The preferred spacer membrane polymer described herein may be used as an outer coating for a diffusion barrier of this disclosure, but also as an outer coating of an electrode system in general, particularly of an electrode system for measuring the concentration of an analyte under in-vivo conditions, comprising an electrode with immobilized enzyme molecules and a diffusion barrier that controls diffusion of the analyte from the exterior of the electrode system to the enzyme molecules.

[0143] Thus, this disclosure provides an electrode system for measuring the concentration of an analyte under in-vivo conditions, comprising an electrode with immobilized enzyme molecules and preferably a diffusion barrier that controls diffusion of the analyte from the exterior of the electrode system to the enzyme molecules, characterized in that a spacer membrane forms at least a portion of the outer layer of the electrode system, wherein the spacer membrane comprises a hydrophilic copolymer of acrylic and/or methacrylic monomers, wherein the polymer comprises more than 50 mol-% hydrophilic monomers.

[0144] The features of this embodiment particularly with regard to the structure of the electrode system, the analyte and the enzyme molecules are as described herein. The diffusion barrier is preferably based on acrylic and/or methacrylic monomers as described herein, it may however also have a different composition.

[0145] The outer spacer membrane preferably covers at least the working electrode portion comprising the enzyme molecules and optionally also other portions, e.g., the counter electrode.

EXAMPLE 1

Permeability of an Enzymatic Non-fluidic (ENF) Glucose Sensor with Distributed Electrodes for Transcutaneous Implantation having a Diffusion Layer Consisting of One Single Block Copolymer

[0146] The sensor was built on a prefabricated palladium strip conductor structure on a polyester substrate having a thickness of 250 m. Working electrode (WE) and counter electrode (CE) were arranged distributed (as shown in FIGS. 1-2).

[0147] The fields of the CE were overprinted with carbon paste, the rest of the strip conductor was insulated. The fields of the WE were overprinted with a mixture of cross-linked glucose oxidase (enzyme), conductive polymer paste and electric catalyst, here manganese (IV)-oxide (Technipur). The remaining paths of the strip conductor were again insulated. The reference electrode (RE) consists of Ag/AgCl paste. The electrodes cover about 1 cm of the sensor shaft.

[0148] The WE-fields were coated with a block copolymer diffusion layer consisting of a HEMA block and a n-BuA block. The thickness of the layer is between 20 and 50 m.

[0149] Three sensor batches were produced, each coated with a specific block copolymer as diffusion layer (see Table 2 hereinbelow). All block copolymers were obtained from Polymer Source, Montreal and TU Darmstadt and are listed in the following Table 2.

TABLE-US-00002 TABLE 2 Number average Name Monomeric units Molecular weight (Mn) block Molar ratio/% nBuA HEMA Copolymer copolymer nBuA/HEMA [n] [n] [kDa] A 45/55 220 270 63 B 15/85 90 522 79 C 17/83 40 190 30

[0150] The respective block copolymer was dissolved in organic solvent (15-18% concentration) and the sensors were coated therewith. After drying in an oven for 1 hour at 50 C. under atmosphere conditions, the coated sensors were tested in-vitro in glucose solutions of different concentrations. Of each sensor batch 4 sensors were measured as random sample. As a measure for the in-vitro sensitivity, the signal was calculated by the difference of the measured currents at 10 mM and 0 mM glucose concentration, which then was divided by 10 mM (cf. Example 4).

[0151] All sensors were operated at a polarisation voltage of 350 mV versus Ag/AgCl, the measurement temperature was kept constant at 37 C. The sensors used for this measurement series did not comprise the spacer described in WO 2010/028708, which, however, did not make any difference in view of the tested signal level. FIGS. 5a and b show the diffusion coefficient and the sensor drift of the sensors with standard deviations for the three different diffusion layers.

[0152] Block copolymer A represents a comparative block copolymer having a molar ratio of 45:55 (hydrophobic:hydrophilic). It shows a high diffusion coefficient associated with a comparatively high sensor drift (as shown in FIGS. 5a and 5b).

[0153] In block copolymer B according to this disclosure, the amount of the hydrophilic HEMA block was increased to 85 mol-% and the hydrophobic nBuA block was reduced to 15 mol-%. Block copolymer B shows a somewhat reduced diffusion coefficient associated with a significantly reduced sensor drift (as shown in FIGS. 5a and 5b).

[0154] Block copolymer C essentially corresponds to block copolymer B with respect to the hydrophobic and hydrophilic portions and molar ratios. However, block copolymer C shows a reduced total number average molecular weight (Mn) compared to block copolymer B. This leads to a further reduction of the diffusion coefficient and to a very low, slightly positive, sensor drift (as shown in FIGS. 5a and 5b).

[0155] Thus, according to this disclosure, a desired diffusion coefficient for the analyte, e.g., glucose within a broad range can be adjusted. Thereby, a preferred sensitivity range of 1 to 1.5 nA/mM analyte, e.g., glucose may be obtained.

EXAMPLE 2

Mechanic Flexibility of the Diffusion Layer of an ENF Glucose Sensor

[0156] The sensor was manufactured as described in WO 2010/028708, however having a diffusion layer according to this disclosure. It was assumed that the glass transition temperature (Tg) is a substitute parameter for the mechanic flexibility. Further, it was assumed that the glass transition temperature, which may be allocated to the hydrophobic block, determines the mechanic flexibility in in-vivo applications.

[0157] The sensors were coated with the diffusion layer of the block copolymer B of this disclosure as described in Example 1 or with a comparative block copolymer Poly(BUMA-r-MMA-b-HEMA) as described in WO 2012/130841 in example 4. The molar ratio of the hydrophobic BUMA-r-MMA block to the hydrophilic HEMA block is 50%:50%, the hydrophobic block contains MMA and BUMA in equal molar amounts in a randomized sequence. The molecular weight is 50 kDa. The glass transition temperature of the randomized hydrophobic block is about 73 C., determined by DSC and a heating rate of 10 C/min.

[0158] Both diffusion layers were generated from the respective solution of the copolymers in organic solvent and dried as in Example 1. The thickness of the diffusion layers was approx. 20 m.

[0159] Bending tests on both sensor types with a bending radius of 1 mm were performed by hand. The sensors having a block copolymer B diffusion layer showed no cracks in the diffusion layer. In contrast thereto, under identical test conditions, the sensors having the block copolymer Poly(BUMA-r-MMA-b-HEMA) showed Newton's rings due to a delamination of the diffusion layer (as shown in FIG. 6).

EXAMPLE 3

Hydrophobicity of the Diffusion Layer of an ENF EC Sensor

[0160] The sensors were manufactured as described in WO 2010/028708 with comparative block copolymer D and block copolymer E of this disclosure comprising hydrophilic blocks of HEMA and hydrophobic blocks of Ethyl acrylate (EA) (c.f. Table 3). All test conditions were identical to Example 1.

TABLE-US-00003 TABLE 3 Name Monomers Number average block Molar ratio/% EA HEMA Molecular weight copolymer EA/HEMA [n] [n] Mn [kDa] D 43/57 290 260 63 E 22/78 100 350 56

[0161] The sensors coated with the block copolymer D were found to be unstable with regard to the sensor behavior. Thus, no value for the diffusion coefficient or the sensor drift could be measured.

[0162] Instead, a good sensor behavior was observed with block copolymer E, wherein the hydrophilic HEMA fraction was increased to about 80 mol% and the hydrophobic EA fraction was decreased to about 20 mol%. As can be seen in FIGS. 7a and 7b, the block copolymer E shows an increased diffusion coefficient, but also an increased sensor drift as compared to block copolymer B from Example 1. This may be due to the lower hydrophobicity of PolyEA compared to Poly(n-BuA).

[0163] A different behavior of block copolymers B and E was also observed at the start of the measurement. Block copolymer E (Poly(EA-b-HEMA)) was found to have a comparatively long-lasting (24 h) start-up phase as shown in FIG. 8b whereas block copolymer B (Poly(nBuA-b-HEMA)) only shows a very short start-up phase as shown in FIG. 8a.

EXAMPLE 4

Characterization of the Block Copolymer

[0164] A multiple field sensor (10 fields of working electrodes and counter electrodes, respectively) for the continuous measurement of the glucose was produced and characterized in-vitro.

[0165] The sensor was provided with a diffusion layer consisting of a block copolymer comprising a hydrophobic block of n-butyl acrylate (n-BuA) and a hydrophilic block of 2-hydroxyethyl methacrylate (HEMA). These polymers correspond to block copolymers B and C of Example 1.

[0166] The decisive parameter with regard to the permeability of the diffusion barrier for the analyte is the sensitivity per area unit of the working electrode (i.e., the geometric area). The sensitivity SE is calculated from current (I) measurements at 10 mM and at 0 mM glucose concentration in phosphate-buffered solution (pH 7.4) in nA/mM:

SE=[i(10 mM)I(0 mM)]10

for each of the analysed sensors. From the individual measurement values (N=8) the mean sensitivity SE.sub.m is determined. The obtained sensitivity values are divided by the microscopically measured geometric total area F of all working electrode spots on the multi-field sensor. Thereby, a sensitivity density SE.sub.m/F is obtained.

[0167] The linearity Y of the in-vitro function curve is an indication of the diffusion control functionality of the polymer cover layer on the working electrode. It is calculated from current measurements at 20 mM, 10 mM and 0 mM glucose concentration in %:

Y.sup.20mM=50.Math.[I(20mM)I(0mM)]/[I(10mM)I(0mM)]

for each of the analysed sensors. From the individual measurement values the mean linearity value and its standard deviation are determined.

[0168] Finally, the layer thickness L of the diffusion barrier of the sensors is determined by optical measurement for each of the polymers. The corresponding mean values are computed for a sample of n=4 sensors with the same polymer. Therefrom, the effective diffusion coefficient D.sub.eff of the cover layer may be calculated:

D.sub.eff=SE.sub.m/F.Math.L.sub.m.Math.5.182.Math.10.sup.8

in cm.sup.2/s, wherein SE.sub.m and L.sub.m are the respective mean values for the sensitivity and the layer thickness, and F is the total area of all working electrode spots.

[0169] The sensor drift was calculated from repetitions of the glucose concentration stages over 7 days of in-vitro measurements.

[0170] The results are shown in Table 4:

TABLE-US-00004 TABLE 4 Parameter A B C Mn Poly(nBuA-b-HEMA) [kDa] 28-b-35 12-b-68 5-b-25 Tg 1 [ C.] 44 49 n.m. Tg 2 [ C.] 115 118 n.m. nBuA [mol %] 45 15 17 HEMA [mol %] 55 85 83 nBuA-molecules [n] 218 90 39 HEMA-molecules [n] 269 522 188 layer thickness L.sub.m [m] 35 47 22 SE.sub.m/F [nA/(mM*mm.sup.2)] 6.5 2.7 0.3 D.sub.eff [cm.sup.2/s] 1.2E08 6.7E09 3.6E10 Drift [%] 3.5 1.1 0.6 Y20 mM[0%] 74 88 83 n.m. = not measured

[0171] While exemplary embodiments have been disclosed hereinabove, the present invention is not limited to the disclosed embodiments. Instead, this application is intended to cover any variations, uses, or adaptations of this disclosure using its general principles. Further, this application is intended to cover such departures from the present disclosure as come within known or customary practice in the art to which this invention pertains and which fall within the limits of the appended claims.