HK1130088B

HK1130088B - Amperometric sensor and method for its manufacturing

Info

Publication number: HK1130088B
Application number: HK09107987.1A
Authority: HK
Inventors: Arnulf Staib; Reinhold Mischler; Martin Hajnsek; Harvey Buck; Walter Jernigan
Original assignee: F. Hoffmann-La Roche Ag
Priority date: 2006-06-19
Filing date: 2007-05-24
Publication date: 2017-08-18

Description

Amperometric sensor and method for manufacturing same

The present invention relates to an amperometric sensor configured for implantation in the living body of a human or animal for measuring the concentration of an analyte in a bodily fluid, the sensor comprising a counter electrode and a working electrode, the working electrode comprising a sensing layer permeable to water and arranged on a contact pad, the sensing layer comprising an immobilized enzyme capable of catalyzing in the presence of the analyte to cause an electrical signal, the sensing layer having a lower surface facing the contact pad and an upper surface facing away from the contact pad. Such sensors are known from EP0247850B 1.

Implantable sensors for in vivo measurement of medically important analytes, such as glucose or lactate, are based on electrochemical enzymatic detection of the analyte. The most common method is to oxidize an analyte such as glucose using an oxidase enzyme, then reduce oxygen to hydrogen peroxide, and amperometrically detect the hydrogen peroxide with the working electrode of the sensor. Another approach in the field of in vivo sensing is to circumvent the use of oxygen/peroxide as a mediator pair by carrying out glucose conversion without oxygen using synthetic redox mediators. In this case, the synthesized redox mediator is embedded within the sensing element. An example of such a method utilizes poly (biimidazolyl) osmium complexes as redox mediators for binding to enzymes, as described by Feldmann et al in diabetes technology and therapeutics, 5, 769 (2003).

Despite extensive research and development efforts, there is currently no implantable sensor that can reliably measure medically important analytes, such as glucose, over long periods of time.

It is an object of the present invention to provide a method for improving the reliability and lifetime of amperometric sensors (amperometric sensors) for in vivo measurement of analyte concentrations in body fluids.

According to the invention, this object is achieved by an amperometric sensor configured for implantation in the living body of a human or animal for measuring the concentration of an analyte in a bodily fluid, said sensor comprising a counter electrode and a working electrode, said working electrode comprising a sensing layer permeable to water and arranged on a support member close to a contact pad, said sensing layer comprising an immobilized enzyme capable of catalyzing in the presence of the analyte to give an electrical signal, the sensing layer having an upper surface facing the bodily fluid and a lower surface facing away from the bodily fluid, characterized in that the immobilized enzyme is distributed in the sensing layer such that the concentration of the enzyme in a middle portion between the upper and lower surfaces of the sensing layer is at least as high as on the upper surface of the sensing layer.

In a planar configuration, the contact pad may be placed directly under the sensing layer (or the sensing layer is disposed on the contact pad), both having the same surface area. In another embodiment, the contact pad may be made smaller or larger than the sensing layer. In yet another embodiment, the contact pad may be partially displaced from the area covered by the sensing layer, so that only a portion of the sensing layer directly contacts the pad. For other arrangements, the contact pad may be placed on one of the sides of the sensing layer. All these alternatives are outlined with the term "sensing layer close to the contact pad". It should be appreciated that the same is true for the other electrodes.

It was found that the measurements of implanted amperometric sensors are often negatively affected by the low oxygen concentration in the subcutaneous tissue surrounding the sensor. This problem appears to be particularly pronounced in the case of enzyme sensors that rely on an oxidase, such as glucose oxidase, as the immobilized enzyme in the sensing layer, since such sensors cause an electrical measurement signal by oxidizing the analyte. In principle, the intensity of the measurement signal caused by such sensors depends on the amount of enzyme, analyte and oxygen present. If the oxygen concentration is sufficiently high, the response of a given sensor with a specified enzyme loading reflects, and ideally is proportional to, the concentration of the analyte in the vicinity of the sensor. However, if the oxygen concentration is too low, fewer analyte molecules are oxidized and, therefore, a weaker electrical signal is generated than a sensor operating under oxygen saturation conditions.

Lowering the enzyme loading of the sensor lowers the critical oxygen concentration to saturation, but also lowers the signal-to-noise ratio because a smaller measurement signal is formed. Therefore, lowering the enzyme loading is not sufficient to solve the problem.

The amperometric sensor according to the invention solves the problem of low oxygen concentration in the subcutaneous tissue with a sensing layer comprising immobilized enzyme distributed in the sensing layer such that the concentration of enzyme in the middle between the upper and lower surfaces is at least as high as on the upper surface of the sensing layer.

Thus, only a relatively small fraction of the enzyme molecules comprised in the sensing layer are active on the upper surface of the sensing layer. Thus, a relatively low oxygen concentration is sufficient to saturate the sensing layer surface with oxygen. The structure of the sensing layer allows analyte molecules to diffuse into the sensing layer and interact with enzyme molecules surrounded by their own pool of oxygen molecules away from the surface. The electrical signal of the sensor according to the invention is thus formed not only in a small surface layer but also in an enlarged volume which reduces the oxygen density (oxygen concentration) which achieves saturation of the sensor. Thus, oxygen saturation of the enzyme can be achieved at lower oxygen concentrations without reducing the signal-to-noise ratio of the sensor measurement signal.

Amperometric sensors with a porous sensing layer have been described in EP0247850B 1. However, the enzyme is only supplied to the known sensor after the porous layer has been prepared. Thus, the enzyme concentration of the sensor described in EP0247850B1 is highest on the upper surface of the sensing layer and decreases sharply with increasing distance from the surface. Thus, a large part of the electrical signal of such sensors is formed on this surface of the sensing layer, i.e. in a relatively small volume, so that a correspondingly higher oxygen concentration is required for an accurate measurement.

The distribution, especially the uniform concentration, of the enzyme over the entire sensing layer according to the invention can very easily be achieved by mixing the enzyme into a paste, preferably a paste comprising carbon particles and a binder, and applying this mixture to the contact pad to provide the sensing layer of the working electrode. In some cases, it may be advantageous to use a surfactant such as a detergent or a hydrophilic polymer to aid in the dispersion of the enzyme within the paste. In this way, an equal distribution of enzyme molecules throughout the sensing layer can be achieved. The object of the invention is thus also achieved by a method of manufacturing an amperometric sensor configured for implantation into the living body of a human or animal for measuring the concentration of an analyte in a body fluid, the method comprising the steps of: mixing carbon particles, an enzyme and a polymeric binder to form a paste; the paste is applied to the support member adjacent the contact pads and hardened into a porous sensing layer.

Further details and advantages of the invention are described below on the basis of exemplary embodiments with reference to the drawings. The features described herein can be used alone or in combination to limit the invention. In the figure:

fig. 1 shows a first exemplary embodiment of a sensor according to the invention in cross-section.

Fig. 2 shows the functional characteristics of the sensor according to fig. 1 from an in vitro measurement.

Fig. 3 shows measurement data measured in a biological matrix by the sensor according to fig. 1.

Fig. 4 shows the dependence of the sensor current on the diffusion barrier covering the sensing layers of sensors F to J instead of sensors a to E.

Fig. 5 shows a second exemplary embodiment of a sensor according to the invention in cross-section.

Fig. 1 schematically shows a first embodiment of an amperometric sensor 1 configured for implantation into the living body of a human or animal to measure the concentration of an analyte in a body fluid of the human or animal. For better illustration of some details, FIG. 1 is not drawn to scale.

The sensor 1 comprises a counter electrode 2, a working electrode 3 and a reference electrode 4 arranged on a support member 5 made of a plastic material, in particular polyimide. Each electrode 2, 3, 4 comprises a contact pad 6, 7, 8 provided as an electrically conductive film, for example a metal film, in particular a gold film, preferably 50nm to 150nm thick. The contact pads 6, 7, 8 can also be made of other metals, in particular palladium, or in the form of multilayer films of different metals. For example, a thin film of titanium less than 20nm covering the support member 5 may be covered by a second gold film having a thickness of 50-130nm, thereby forming the contact pads 6, 7, 8. Alternatively, the contact pads 6, 7, 8 may be formed in the form of a conductive polymer film, for example from a conductive polymer paste by, for example, screen printing or by dispensing resulting in thicker contact pads 6, 7, 8. Instead of separate counter and reference electrodes 3, 4, a combined counter/reference electrode may also be used. One example of a suitable counter/reference electrode is a silver/silver chloride electrode. Since such counter and/or reference electrodes are commonly used, further description is not required.

The working electrode 3 also includes a sensing layer 9 that is permeable to water and is disposed proximate to the contact pad 7 of the working electrode 3. The sensing layer 9 comprises an immobilized enzyme capable of catalyzing the action of an analyte to cause an electrical signal. In the present example, an oxidase, in particular glucose oxidase, is used as an enzyme for measuring glucose as an analyte in a human body fluid such as interstitial fluid or blood.

The sensing layer 9 is applied in the form of a paste to the support member 5 over the contact pad 7 covering the working electrode 2. The paste is prepared by mixing carbon particles, an enzyme and a polymeric binder. In this way, the immobilized enzyme is equally distributed throughout the sensing layer 9. A homogeneous enzyme distribution throughout the sensing layer 9 is advantageous. Thus, the enzyme concentration should differ by less than 20%, in particular less than 10%, between the upper and lower surface of the sensor layer 9. Since the analyte can diffuse into the interior of the porous sensing layer 9, an electrical measurement signal is generated not only in the upper surface of the sensing layer 9 facing away from the contact pad 7, but also within the extended volume. Thus, a rather low oxygen concentration is sufficient to saturate the sensor 1 with oxygen and to enable accurate measurements.

In a preferred embodiment, the sensing layer 9 is flat. Preferably, the sensor layer 9 is electrically conductive, wherein the electrical conductivity of the sensor layer 9 is at least 1 Ω^-1cm^-1. This achieves the advantage that each location in the sensor layer 9 where an enzymatic reaction of the analyte takes place serves as a microelectrode at which the products of the enzymatic reaction can be reduced or oxidized directly. In this way, these sites act as cathodes or anodes depending on the applied potential signal. The sensing layer 9 in the porous structure thus comprises a large number of tiny cathodes or anodes. As a result, products that do not require enzymatic reactions proceed through the entirety of the sensing layer 9, which passage can result in a loss of signal height. The conductive embodiment of the sensing layer 9 thus has an increased signal height.

The sensing layer 9 of the shown example has a thickness of 30 μm. In general, the sensor layer 9 should have a thickness of at least 5 μm, preferably at least 10 μm, in order to provide a sufficiently large volume for the formation of electrical measurement signals. A thickness of the sensing layer 9 exceeding 100 μm does not provide additional benefits. A thickness of 20 μm to 70 μm is preferred. The sensing layer 9 is arranged in a recess of the support member 5. In this way it is protected to some extent by the support member 5 side walls from damage during implantation. Furthermore, the side surfaces of the sensing layer 9 may be connected to the support member 5 and thus ensure that analyte molecules may diffuse into the sensing layer 9 only through the upper surface of the sensing layer 9. Of course, the side surfaces can also be made watertight by different means. The sensing layer 9 may have a side surface that is impervious to bodily fluids.

In a similar manner, the contact pads 6, 8 of the counter electrode 2 and the reference electrode 4 are covered by water permeable layers 12, 14, which are also applied in paste form. Of course, the layers 12, 14 of the counter electrode 2 and the reference electrode 4 do not contain enzymes. Like the sensing layer 9, the layers 12 and 14 may also include carbon particles and a polymer binder. Although porosity enhancing particles 13, such as carbon nanotubes, have been added to the paste for the sensing layer 9 and the layer 12, such porosity enhancing particles 13 do not provide a benefit to the strongly conductive layer 14 of the reference electrode 4 and therefore are not added.

Since the enzyme is distributed throughout the sensing layer 9, oxygen saturation is maintained even if a much higher analyte concentration is present at the upper surface of the sensing layer 9 than is feasible with known sensors. The sensing layer of sensors according to the prior art is usually covered by a diffusion barrier which hinders the diffusion of the analyte to such an extent that the analyte concentration at the upper surface is typically about 100 times lower than the analyte concentration in the body fluid surrounding the sensor.

The sensing layer 9 of the sensor 1 of the present embodiment is covered by a diffusion barrier which hinders diffusion of analyte molecules only to such an extent that after implantation into the living body of a human or animal the analyte concentration at the upper surface of the sensing layer 9 is at least one tenth, in particular at least one fifth, preferably at least one third of the body fluid surrounding the implanted sensor 1. In the example shown, the diffusion barrier comprises a plurality of different layers 10, 11 that contribute to the diffusion resistance of the diffusion barrier to diffusion of analyte molecules.

The diffusion barrier allows analyte to permeate and prevents enzyme from leaking out of the sensing layer 9. In the example shown, the diffusion barrier comprises as a first layer an electrically conductive enzyme-free layer 10 comprising carbon particles and a polymer binder and having a thickness of less than one third of the thickness of the sensing layer 9. Typically it is about 1 μm to 3 μm thick. Like the sensing layer 9, the enzyme-free layer 10 is applied in the form of a paste. This paste differs from the paste of the sensing layer 9 only in that no enzyme is added thereto.

The diffusion barrier further comprises a layer 11 that prevents macromolecules from blocking the pores of the sensing layer 9. The layer 11 may be a dialysis layer provided in the form of a membrane made of cellulose and/or polymer material. Such dialysis layers are also enzyme-free layers and can be applied directly on top of the sensing layer 9 or, as shown in fig. 1, on top of the electrically conductive enzyme-free layer 10. It is advantageous if such a dialysis layer impedes the diffusion of the analyte as little as possible. Preferably, the layer 11 has an effective diffusion coefficient D of the analyte which is at least 10 times the diffusion coefficient D of the analyte in water_effIn particular at least 5 times the diffusion coefficient D of the analyte in water. The dialysis layer may be applied as a solid membrane or as a polymer solution that hardens into a dialysis membrane in situ.

Dialysis membranes are generally characterized by their molecular weight cut-off (MWCO), which depends on the pore size. MWCO describes the molecular weight of a compound that retained 90% after overnight (17-hour) dialysis. The dialysis layers of the illustrated examples have an MWCO of less than 10kD (kDaltan), preferably less than 7kD, especially less than 5 kD. It must be recognized that the MWCO specified for the dialysis layer is strictly applicable only to globular molecules such as most proteins. More linear molecules can pass through the pores of the dialysis layer even if their molecular weight exceeds the indicated MWCO.

Instead of or in addition to the dialysis membrane, the diffusion barrier may also comprise a polymer layer made of a polymer having a zwitterionic structure to protect the sensing layer 9 and any porous layer 10 from protein ingress. The zwitterionic structure allows for rapid absorption of polar protic solvents, especially water, and analytes such as internally dissolved glucose. Thus, polymers having zwitterionic structures attached to a polymer backbone are not permeable to proteins, but very slightly impede the diffusion of analytes such as glucose. A well-known example of such a polymer is poly (2-methacryloyloxyethyl phosphorylcholine-co-n-butyl methacrylate) (abbreviated MPC). The MPC polymer layer 11 is applied in the form of a polymer solution comprising ethanol or distilled water and at least 5 wt% MPC, especially at least 10 wt% MPC.

The diffusion barrier and in particular the polymer layer 11 it comprises protects the sensor 1 from mechanical damage during implantation, prevents leakage of enzymes out of the sensing layer 9 into the surrounding tissue where enzymes may be harmful, and prevents macromolecules from blocking the pores of the sensing layer 9. A polymer having a zwitterionic structure, such as MPC, and another polymer, such as polyurethane, or a typical component of the dialysis membrane described above, may be mixed in order to adjust the physical properties of the polymer layer 11.

If layer 11 comprises a copolymer of components having different hydrophilicity, then the physical properties of layer 11, such as permeability to an analyte, can also be adjusted by varying the relative amounts of each component in the copolymer. In the case of MPC, the relative amount of 2-methacryloyloxyethyl phosphorylcholine p-butyl methacrylate can be increased from 30: 70% to 50: 50% to produce a copolymer with higher permeability to polar protic solvents or glucose. Another method to increase permeability to polar protic solvents or glucose is to change the hydrophobic backbone of the copolymer to a more hydrophilic entity. This also applies to other water-soluble analytes.

The sensing layer 9 of the example shown in fig. 1 comprises porous particles 13 to increase its porosity and thus facilitate diffusion of analyte molecules into the sensing layer 9. The porous particles 13 in this aspect are particles having voids that adsorb water molecules. These porous particles 13 are added to the paste forming the sensing layer 9 and create voids through which analyte molecules and water can pass. The porous particles 13 are bound to the other particles of the paste by a polymeric binder. Carbon nanotubes are particularly useful additives to increase the porosity of the sensing layer because they tend to form coils, which are only partially filled with carbon particles and binder, and also increase the electrical conductivity of the sensing layer. Silica particles may also be used as porous particles 13 to increase the porosity of the sensing layer 9.

If silica or similar porous particles 13 are used, it is advantageous to use a material having a particle size distribution such that the maximum particle size is smaller than the thickness of the sensing layer 9. For being most effective, the porous particles 13 should have a size of at least 1 μm, in particular at least 5 μm. Considering a sensing layer 9 thickness of about 20 μm to 50 μm, the silica FK320 from Degussa provides sufficient particle size of up to 15 μm. Typically, less than 10% of such material is mixed into the paste, preferably less than 5%.

It is important to provide electrical conductivity throughout the sensing layer 9 so that at each point of the porous matrix where a product molecule is produced by an enzymatic reaction, such molecule is directly oxidized or reduced by application of a suitable voltage without the need for such molecule to spread to a remote location. In these cases, the porous permeable sensing layer 9 is capable of electrolyzing the analyte over substantially the entire layer.

Regardless of the means used to increase porosity, the mixing of the enzyme with the paste will result in a portion of the enzyme molecules being accessible to the analyte, either on the upper surface of the sensing layer 9 or at the channels near the additive particles within the sensing layer 9. The enzyme is immobilized by adsorption and entrapment in the working electrode 3. The entrainment depends not only on the sensing layer 9 but also on the properties of the diffusion barrier layer, i.e. layer 11, and optionally the properties of the enzyme-free layer 10. It will be appreciated that in order to maintain the desired distribution of the enzyme within the working electrode, contact with the solvent (water) should not result in substantial detachment of the enzyme from the matrix and subsequent migration of the enzyme molecules. The enzyme immobilization in the sensing layer 9 may be enhanced by cross-linking. Particularly advantageous are enzyme molecules which are cross-linked as chains. If these chains are too long, the enzyme is less effective. It is therefore preferred that on average 3 to 10, especially 4 to 8, enzyme molecules are linked together. Chain lengths of 5 to 7 enzyme molecules appear to be most advantageous.

A cross-linking agent such as a glutaraldehyde solution may be added to the paste prior to drying. However, it is preferred to mix already cross-linked enzymes into the paste. It is advantageous to use enzymes which form complexes with hydrophilic partners. After mixing into a less hydrophilic or even hydrophobic paste, as can be achieved by mixing the carbon particles with a suitable binder, the cross-linked enzyme is located in a local hydrophilic environment that contributes to its stability. An additional advantage of cross-linked enzymes with hydrophilic partners is that it enhances the migration of hydrated analyte molecules towards the enzyme. Thus accelerating the wetting of the sensing layer 9, which shortens the wetting time of the sensor after implantation. In particular, for example, glucose oxidase crosslinked with dextran from Roche diagnostics (Penzberg, Germany, Ident-No.1485938001) has been found to have such an enzyme content (approximately 16%) that sufficient activity (20-40U/mg of lyophilizate) can be maintained. Due to the high degree of hydrophilic dextran in the complex, the above advantages can be exploited.

By mixing the already cross-linked enzyme with the sensing layer paste comprising carbon nanotubes, the character of the carbon nanotubes intertwining and forming coils is supported by the larger enzyme-dextran chains, in particular by their aggregation, wherein the coils act as a macroporous cage structure. Thus, the cross-linking enzyme will contribute to the formation of the porous structure of the sensing layer 9.

The sensing layer 9 of the shown example comprises carbon particles having an average size of less than 1 μm, a polymer binder, an enzyme and carbon nanotubes as porous particles 13. The porous particles 13 most effectively increase the porosity of the sensing layer 9 if they are significantly larger than the carbon particles. In the example shown, the porous particles 13 have on average a size of at least 1 μm, in particular at least 5 μm. Typically, the sensing layer 9 comprises 50 to 70 wt% of a polymeric binder, 20 to 40 wt% of carbon particles and not more than about 20 wt%, preferably 1 to 10 wt% of porous particles 13 such as carbon nanotubes or silica. Carbon nanotubes are particularly advantageous additives because they increase both the porosity of the sensing layer 9 and the electrical conductivity of the sensing layer 9. In the embodiment schematically shown in FIG. 1, multiwall carbon nanotubes of NanoLab, Newton, MA (research grade, > 95% purity) are used, having a length of 5 μm to 20 μm and an average outer diameter of 25nm to 35 nm. The binder is a thermoplastic resin, for example based on epoxy resin or on polyvinyl chloride (PVC)/polyvinyl alcohol (PVA). It is also possible to use as binder a resin based on a fluorocarbon resin, in particular polytetrafluoroethylene or based on polystyrene. In the case of PVC/PVA binders, the use of additives such as silicone oils can help to adjust the viscosity of the paste.

In this way, the sensing layer 9 of the sensor 1 shown in fig. 1 is adapted and arranged such that in post-implantation operation the analyte concentration in the sensing layer 9 is highest at the upper surface and decreases with increasing distance from the upper surface and is zero at the lower surface which is the furthest point from the analyte-containing body fluid and which contacts the contact pad 7. The enzyme loading of the sensing layer 9, i.e. the amount of enzyme immobilized therein, should be selected with respect to the porosity and water permeability of the sensing layer 9.

An important parameter in this respect is the effective diffusion coefficient D of the sensing layer 9_eff. Effective diffusion coefficient D_effThe diffusion of the analyte in the sensing layer 9 is characterized and depends on the pore volume and the curvature τ of the sensing layer 9. In general, the effective diffusion coefficient D_effCan be described as D_effD ·/τ, where D is the diffusion coefficient of the analyte in water. The quotient τ/is also referred to as the hindrance H. In the example shown, H is between 10 and 1000, in particular between 50 and 500.

Another important parameter in this aspect is the enzyme loading parameter α, which can be described as α ═ V (V)_max·d)/(K_MD) where V_maxEnzyme Activity Density to determine maximum Rate of analyte conversion, K_MIs the Michaelis Menten constant of the enzyme, D is the thickness of the sensing layer, and D is the diffusion coefficient of the analyte in water. The effective diffusion coefficient D in the sensing layer 9 is preferred_effThe ratio of the enzyme loading parameter α was in the range of 10-200.

Figure 2 shows the functional characteristics of the sensor described above. Layer 11 was made from MPC (lipidure cm5206, NOFCorp, japan) by dispensing a 10% solution of MPC in ethanol/water onto an electrode. The measured current I in nA is plotted against the glucose concentration g in mg/dl. The data shown in figure 2 were measured in vitro in aqueous glucose solution. As can be seen, no saturation was observed at the higher glucose concentration.

FIG. 3 shows the measured current I in nA for comparison_AAnd I_BCurrent I of_AMeasured in vitro, current I_BMeasurements were made with the sensors in a bio-matrix, both after the sensors had equilibrated in the respective medium for 12 hours at a temperature T ═ 35 ℃. Each data point displayed thus pertains to a bio-matrix measurement and an aqueous glucose solution measurement at the same glucose concentration. The biological matrix used consists of stabilized plasma to which glucose is added to obtain the desired glucose concentration. The sensor current measured in the biological matrix and the sensor current measured in the aqueous glucose solution showed excellent consistency.

The results are particularly noteworthy and demonstrate the significant effect of the sensor arrangement resulting from embodiments of the present invention. In general, it is expected that exposure of the sensor 1 to the biological matrix ensures that proteins, peptides or fibrin are deposited on the sensor surface. This process affects the permeability of the outer layer, such as layer 11, to the analyte or water. In conventional sensor arrangements, this layer limits the diffusion of analyte to the sensing layer, so that a decrease in permeability results in a weaker measurement signal.

However, the signal height of the sensor 1 is not affected by exposure of the bio-matrix, as shown in fig. 3, since diffusion of analyte through the layer 11 is not a rate limiting step in signal generation. Thus, any change in permeability has little effect on the signal of the sensor 1 described above.

This advantage of the present invention is not limited to enzymes that use oxygen as a co-substrate in a catalytic reaction. The enzyme may also be a dehydrogenase. For example, a glucose dehydrogenase that does not use oxygen as a co-substrate may be distributed within the sensing layer 9. Known dehydrogenases include certain molecules as cofactors for glucose oxidation, such as pyrroloquinoline quinone (PQQ) or Flavin Adenine Dinucleotide (FAD) or Nicotinamide Adenine Dinucleotide (NAD), see EP1661516a 1. Any of these dehydrogenases may be used in the sensing layer 9 instead of the oxidase.

Fig. 4 shows the sensor current I in nA, measured with sensors a to J in phosphate buffered aqueous glucose solutions of different concentrations. The sensor current measured at a glucose concentration of 360mg/dl is depicted by triangles (a). The sensor current measured at a glucose concentration of 180mg/dl is depicted by a square (■). The sensor current at zero glucose concentration is depicted in diamond shape (° c).

The sensors a to J differ only in the diffusion barrier layer applied on top of the sensing layer 9. In sensors a to E, the diffusion barrier is absent, i.e. the sensing layer 9 is in direct contact with the aqueous glucose solution to be measured. The sensors F to J include diffusion barriers covering the sensing layer 9. The diffusion barrier of the sensors F to J is provided in the form of a polymer-layered layer 11 made of MPC in fig. 1. It can be seen that the sensor currents of sensors F to J are only slightly lower than the sensor currents of sensors a to E. Thus, the diffusion barrier provided by the MPC polymer layer 11 hinders diffusion of analyte molecules only to a small extent. Since the sensor currents of sensors F to J are about 20% lower than the sensor currents of sensors a to E, it can be concluded that the diffusion barrier of sensors F to J results in an analyte concentration on the upper surface of the sensing layer 9 which is only about 20% lower than in the glucose solution surrounding the sensors.

As in the embodiment described previously in connection with FIG. 1, the sensing layer 9 of sensors A to J comprises a cross-linked enzyme, i.e.dextranized glucose oxidase, available as ident-No.14859389001 from Roche diagnostics, Penzberg, Germany. The dextranized glucose oxidase is dissolved in a phosphate buffered solution and mixed into a paste comprising carbon particles, carbon nanotubes and a polymeric binder. A sensing layer 9 is dispensed on the contact pad 7 of the working electrode 3 on the sensor substrate 5 with a spot size of about 0.05mm²To 0.1mm²For example a circular spot of 300 μm diameter. The thickness of the sensing layer 9 was 20 μm. An Ag/AgCl reference electrode 4 having the same dimensions is also provided. The counter electrode 2 has a rectangular shape (400 μm by 900 μm) with a carbon paste layer containing carbon nanotubes 20 μm thick.

As can be seen from fig. 4, the sensor current is hardly affected by the presence of the membrane made of MPC.

From this finding it can be concluded that by specifically selecting the zwitterionic membrane structure, a coating is found that is highly permeable to solvated glucose. This high permeability of the membrane is important for the construction of the sensor 1 when there is a diffusion limitation in the sensing layer 9 (see fig. 1). Vice versa, the diffusion of analyte through the diffusion barrier provided by the MPC layer 11 should be hindered as little as possible, ideally the analyte concentration (i.e. signal) at the sensing layer 9 with the coating should be no less than half the value obtained without the coating.

It should be noted that the optional enzyme-free layer 10 should also have little barrier to analyte diffusion and therefore its layer thickness should be much thinner than that of the sensing layer 9.

As mentioned above, the incorporation of hydrophilic cross-linking enzymes can lead to very stable functions over a long period of time, since the wetting of the sensing layer is fast and the enzyme distribution remains unchanged. This is reflected by the drift value (driftvalue) obtained when measuring the above sensor in an aqueous glucose solution over 6 days. For sensors without coating, the drift ranges from-0.62% per day to 0.78% per day, while those with coating cover the range from-0.5% to 1.5% per day. These small drift values were measured at 37 ℃. The particular advantage of measurement stability, i.e., low signal drift, is not limited to the use of oxygen as a common substrate

The sensor layer 9 of the object has the sensor 1 of the enzyme therein. In fact, the same crosslinking benefits can be obtained by using a crosslinking dehydrogenase that does not require oxygen as a co-substrate in the catalytic reaction. For example, dextranized glucose dehydrogenase or dehydrogenase with polyethylene glycol added (PEG: polyethylene glycol) can be introduced into the sensing layer 9.

In the sensor 1 shown in fig. 1, the sensing layer 9 is arranged on the contact pad 7. In addition, the sensing layer 9 has a lower surface facing the contact pad 7 and an upper surface facing away from the contact pad 7, or more generally, the sensing layer 9 has a lower surface facing the support member 5 and an upper surface facing away from the support member 5 towards the analyte-containing body fluid. Thus, the layers 12, 14 are arranged to the contact pads 6, 8. Fig. 5 shows a modified embodiment of the sensor 1 of fig. 1. The embodiment of fig. 5 corresponds to the embodiment of fig. 1 with the difference that the contact pads 6, 7, 8 of the electrodes are placed on the side of the water permeable layers 9, 12 and 14, in contrast to fig. 1. Contact pads 6, 7, 8 may also be placed on both sides of the respective layers 9, 12, 14, as shown by contact pad 6 being the water permeable layer 12 of the counter electrode 2. The contact pad 6 may also be formed such that it surrounds the layer 12 from all sides. In all cases where the contact pads 6, 7, 8 are located on the side of the permeation layers 9, 12, 14, the surface of the layer 9, 12, 14 facing away from the analyte-containing bodily fluid directly contacts the support member 5.

List of reference numerals

1 sensor

2 counter electrode

3 working electrode

4 reference electrode

5 support member

62 contact pad

73 contact pad

84 contact pad

9 sensing layer

10 enzyme-free layer

11 layers (dialysis layer, Polymer layer, MPC layer)

122 water permeable layer

Porosity enhancing particles (porous particles)

144 of water permeable layer

Claims

1. Amperometric sensor (1) configured for implantation into the living body of a human or animal for measuring the concentration of an analyte in a body fluid, said sensor (1) comprising

A counter electrode (2) and

a working electrode (3) having a plurality of electrodes,

the working electrode (3) comprising a sensing layer (9) permeable to water and arranged on the support member (5) adjacent to the contact pad (7), the sensing layer (9) comprising an immobilized enzyme capable of catalyzing in the presence of the analyte to induce an electrical signal,

the sensing layer (9) has an upper surface facing the body fluid and a lower surface facing away from the body fluid,

is characterized in that

The immobilized enzyme is distributed in the sensing layer (9) such that the enzyme concentration intermediate the upper and lower surfaces is at least as high as on the upper surface of the sensing layer (9),

wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of the analyte in the sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, and

wherein the sensing layer (9) has a thickness of at least 5 μm.

2. The sensor (1) according to claim 1, wherein the sensing layer (9) has a thickness of at least 10 μm.

3. The sensor (1) according to claim 1, wherein the enzyme is a cross-linked enzyme.

4. A sensor (1) according to claim 3, wherein the cross-linked enzyme forms a complex with a hydrophilic partner.

5. Sensor (1) according to claim 4, wherein the cross-linked enzyme is an enzyme wherein the enzyme molecule forms a cross-linked complex with a hydrophilic partner.

6. A sensor (1) according to any of claims 1-5, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme, and wherein the cross-linked enzyme forms a complex with a hydrophilic partner.

7. Sensor (1) according to claim 6, wherein the cross-linked enzyme is an enzyme wherein the enzyme molecule forms a cross-linked complex with a hydrophilic partner.

8. A sensor (1) according to claim 3, wherein the cross-linked enzyme has an average chain length of 3 to 10 enzyme molecules.

9. The sensor (1) according to claim 8, wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules.

10. A sensor (1) according to any one of claims 1-5, 8 and 9, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of said analyte in water, wherein the enzyme is a cross-linked enzyme, and wherein the cross-linked enzyme forms a complex with a hydrophilic partner, and wherein the cross-linked enzyme has an average chain length of 3 to 10 enzyme molecules.

11. The sensor (1) according to claim 10, wherein the cross-linked enzyme is an enzyme wherein the enzyme molecules form cross-linked complexes with hydrophilic partners, and wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules.

12. A sensor (1) according to any one of claims 1-5, 7-9 and 11, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation into a human or animal living body the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymer binder.

13. The sensor (1) according to claim 12, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

14. According to claim1, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the sensing layer (9) is covered by a diffusion barrier layer which hinders diffusion of analyte molecules to such an extent that after implantation in a human or animal living body the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier layer comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

15. The sensor (1) according to claim 12, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the sensing layer (9) is covered by a diffusion barrier layer which hinders diffusion of analyte molecules to such an extent that after implantation in a human or animal living body the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier layer comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

16. The sensor (1) according to claim 14 or 15, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

17. A sensor (1) according to any one of claims 1-3, wherein the enzyme is a cross-linked enzyme, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation into a human or animal living body the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymer binder.

18. The sensor (1) according to claim 17, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

19. A sensor (1) according to claim 12, wherein the enzyme is a cross-linked enzyme, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that, after implantation into the living body of a human or animal, the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

20. The sensor (1) according to claim 19, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

21. A sensor (1) according to any one of claims 1-3, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation into a human or animal living body the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

22. A sensor (1) according to claim 21, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

23. Sensor (1) according to claim 12, wherein the sensor is a piezoelectric sensorThe sensitive layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation into a human or animal living body the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

24. The sensor (1) according to claim 23, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

25. A sensor (1) according to any of claims 1-5, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme, wherein the cross-linked enzyme forms a complex with a hydrophilic partner, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation into the living body of a human or animal the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

26. A sensor (1) according to claim 25, wherein the cross-linked enzyme is an enzyme wherein enzyme molecules form cross-linked complexes with hydrophilic partners, and wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

27. The sensor (1) according to claim 12, wherein the sensing layer (9) has a layer withEffective diffusion coefficient D_effCharacterised by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme, wherein the cross-linked enzyme forms a complex with a hydrophilic partner, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation into the living body of a human or animal the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

28. A sensor (1) according to claim 27, wherein the cross-linked enzyme is an enzyme wherein enzyme molecules form cross-linked complexes with hydrophilic partners, and wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

29. A sensor (1) according to any one of claims 1-5, 8 and 9, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterized by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme, wherein the cross-linked enzyme forms a complex with a hydrophilic partner, wherein the cross-linked enzyme has an average chain length of 3 to 10 enzyme molecules, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation into the living body of a human or animal the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in the body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

30. A sensor (1) according to claim 29, wherein the cross-linked enzyme is an enzyme wherein enzyme molecules form cross-linked complexes with hydrophilic partners, wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules, and wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

31. The sensor (1) according to claim 12, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterized by diffusion of an analyte in a sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme, wherein the cross-linked enzyme forms a complex with a hydrophilic partner, wherein the cross-linked enzyme has an average chain length of 3 to 10 enzyme molecules, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation into the living body of a human or animal the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in the body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder.

32. The sensor (1) according to claim 31, wherein the cross-linked enzyme is an enzyme wherein enzyme molecules form cross-linked complexes with hydrophilic partners, wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules, and wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

33. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28 and 30-32, wherein the sensing layer (9) comprises porous particles (13).

34. The sensor (1) according to claim 33, wherein the porous particles (13) are silica and/or carbon nanotubes.

35. The sensor (1) according to claim 33, wherein the porous particles (13) increase the porosity and the electrical conductivity of the sensing layer (9).

36. A sensor (1) according to claim 33, wherein the porous particles (13) have a size of at least 1 μm on average.

37. A sensor (1) according to claim 36, wherein the porous particles (13) have a size of at least 5 μm on average.

38. The sensor (1) according to claim 33, wherein the sensing layer (9) comprises porous particles (13), and wherein the porous particles (13) have a size of at least 1 μm on average.

39. The sensor (1) according to claim 38, wherein the porous particles (13) are silica and/or carbon nanotubes, and wherein the porous particles (13) have a size of at least 5 μm on average.

40. A sensor (1) according to any one of claims 1, 2, 36 and 37, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in the sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, and wherein the porous particles (13) have a size averaging at least 1 μm.

41. A sensor (1) according to claim 40, wherein the porous particles (13) have a size of at least 5 μm on average.

42. A sensor (1) according to any one of claims 3, 36 and 37, wherein the enzyme is a cross-linked enzyme and wherein the porous particles (13) have a size of at least 1 μm on average.

43. A sensor (1) according to claim 42, wherein the porous particles (13) have a size of at least 5 μm on average.

44. A sensor (1) according to any one of claims 1-3, 36 and 37, wherein the sensing layer (9) has an effective diffusion coefficient D_effCharacterised by diffusion of an analyte in the sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme and wherein the porous particles (13) have a size of at least 1 μm on average.

45. A sensor (1) according to claim 44, wherein the porous particles (13) have a size of at least 5 μm on average.

46. The sensor (1) according to any one of claims 4, 5, 36 and 37, wherein the enzyme is a cross-linked enzyme, wherein the cross-linked enzyme forms a complex with a hydrophilic partner, and wherein the porous particles (13) have a size of at least 1 μm on average.

47. A sensor (1) according to claim 46, wherein the cross-linked enzyme is an enzyme wherein the enzyme molecule forms a cross-linked complex with a hydrophilic partner, and wherein the porous particles (13) have a size of at least 5 μm on average.

48. The sensor (1) according to any one of claims 8, 9, 36 and 37, wherein the enzyme is a cross-linked enzyme, wherein the cross-linked enzyme has an average chain length of 3 to 10 enzyme molecules, and wherein the porous particles (13) have a size of at least 1 μm on average.

49. The sensor (1) according to claim 48, wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules, and wherein the porous particles (13) have a size of at least 5 μm on average.

50. A sensor (1) according to any one of claims 1-5, 8, 9, 36 and 37, wherein the sensing layer (9) has an effective diffusion coefficient D_effWatch, thereofCharacterising the diffusion of the analyte in the sensing layer (9) and being 1/10-1/1000 of the diffusion coefficient D of the analyte in water, wherein the enzyme is a cross-linked enzyme, wherein the cross-linked enzyme forms a complex with a hydrophilic partner, wherein the cross-linked enzyme has an average chain length of 3 to 10 enzyme molecules, and wherein the porous particles (13) have a size of on average at least 1 μm.

51. Sensor (1) according to claim 50, wherein the cross-linked enzyme is an enzyme wherein the enzyme molecules form cross-linked complexes with hydrophilic partners, wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules, and wherein the porous particles (13) have a size of on average at least 5 μm

52. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49 and 51, wherein the sensing layer (9) is adapted and arranged such that in post-implantation operation the analyte concentration in the sensing layer (9) is highest at the upper surface and decreases with increasing distance from the upper surface and is zero at the lower surface, which is the furthest point from the analyte-containing body fluid.

53. Sensor (1) according to any one of claims 1-5, 8, 9, 36 and 37, configured for implantation in the living body of a human or animal for measuring an analyte concentration in a body fluid, the sensor (1) comprising

A counter electrode (2) and

a working electrode (3) having a plurality of electrodes,

is characterized in that

wherein the sensing layer (9) has an effective diffusion coefficient D_effWhich characterizes the diffusion of the analyte in the sensing layer (9) and is 1/10-1/1000 of the diffusion coefficient D of the analyte in water,

wherein the enzyme is a cross-linked enzyme,

wherein the cross-linked enzyme forms a complex with a hydrophilic partner,

wherein the cross-linked enzyme has an average chain length of 3 to 10 enzyme molecules,

wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that after implantation in a living human or animal body the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 of the concentration in a body fluid surrounding the implanted sensor (1), and wherein the diffusion barrier comprises an electrically conductive enzyme-free layer (10) comprising carbon particles and a polymeric binder

Wherein the porous particles (13) have a size of at least 1 μm on average, and

wherein the sensing layer (9) is adapted and arranged such that in post-implantation operation the analyte concentration in the sensing layer (9) is highest at the upper surface and decreases with increasing distance from the upper surface and is zero at the lower surface, which is the furthest point from the analyte-containing body fluid.

54. A sensor (1) according to claim 53, wherein the cross-linked enzyme is an enzyme wherein the enzyme molecules form cross-linked complexes with hydrophilic partners, wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1), and wherein the porous particles (13) have a size of at least 5 μm on average.

55. Sensor (1) according to claim 12, configured for implantation in the living body of a human or animal for measuring an analyte concentration in a body fluid, the sensor (1) comprising

A counter electrode (2) and

a working electrode (3) having a plurality of electrodes,

is characterized in that

wherein the enzyme is a cross-linked enzyme,

wherein the cross-linked enzyme forms a complex with a hydrophilic partner,

Wherein the porous particles (13) have a size of at least 1 μm on average, and

56. The sensor (1) according to claim 55, wherein the cross-linked enzyme is an enzyme wherein enzyme molecules form cross-linked complexes with hydrophilic partners, wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1), and wherein the porous particles (13) have a size of at least 5 μm on average.

57. Sensor (1) according to claim 52, configured for implantation in the living body of a human or animal for measuring an analyte concentration in a body fluid, the sensor (1) comprising

A counter electrode (2) and

a working electrode (3) having a plurality of electrodes,

is characterized in that

wherein the enzyme is a cross-linked enzyme,

wherein the cross-linked enzyme forms a complex with a hydrophilic partner,

Wherein the porous particles (13) have a size of at least 1 μm on average, and

58. A sensor (1) according to claim 57, wherein the cross-linked enzyme is an enzyme wherein the enzyme molecules form cross-linked complexes with hydrophilic partners, wherein the cross-linked enzyme has an average chain length of 4 to 8 enzyme molecules, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1), and wherein the porous particles (13) have a size of at least 5 μm on average.

59. A sensor (1) according to any one of claims 33, 55 and 57, wherein the sensing layer (9) comprises porous particles (13).

60. The sensor (1) according to claim 59, wherein the porous particles (13) are silica and/or carbon nanotubes.

61. The sensor (1) according to claim 53, wherein the sensing layer (9) comprises porous particles (13).

62. A sensor (1) according to claim 61, wherein the porous particles (13) are silica and/or carbon nanotubes.

63. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58 and 60-62, wherein the contact pad (7) of the working electrode (3) is an electrically conductive film.

64. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58 and 60-62, wherein the contact pad (7) of the working electrode (3) is a metal film or a conductive polymer film.

65. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58 and 60-62, wherein the working electrode (3) is arranged on the support member (5).

66. A sensor (1) according to claim 65, wherein the working electrode (3) is arranged on a support member (5) of plastic material.

67. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62 and 66, wherein the sensing layer (9) comprises carbon particles and a polymer binder.

68. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62 and 66, wherein the sensing layer (9) is covered by a diffusion barrier which hinders diffusion of analyte molecules to such an extent that, after implantation in the living body of a human or animal, the analyte concentration at the upper surface of the sensing layer (9) is at least 1/10 in a body fluid surrounding the implanted sensor (1).

69. The sensor (1) according to claim 68, wherein the analyte concentration at the upper surface of the sensing layer (9) is at least 1/5 in a body fluid surrounding the implanted sensor (1).

70. A sensor (1) according to claim 68, wherein the diffusion barrier comprises a dialysis layer (11).

71. The sensor (1) according to claim 68, wherein the diffusion barrier comprises a polymer layer (11) made of a polymer having a zwitterionic structure.

72. The sensor (1) according to claim 69 or 70, wherein the diffusion barrier comprises a polymer layer (11) made of a polymer having a zwitterionic structure.

73. The sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62, 66 and 69-71, wherein the sensing layer (9) is flat.

74. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62, 66 and 69-71, wherein the sensing layer (9) has a lateral surface which is impervious to body fluids.

75. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62, 66 and 69-71, wherein the sensing layer (9) is electrically conductive.

76. A sensor (1) according to claim 75, wherein the sensing layer (9) has at least 1 Ω^-1cm^-1The electrical conductivity of (1).

77. A sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62, 66, 69-71 and 76, wherein the enzyme is an oxidase.

78. The sensor (1) according to claim 77, wherein the enzyme is glucose oxidase.

79. The sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62, 66, 69-71, 76 and 78, wherein the enzyme is a dehydrogenase.

80. The sensor (1) according to claim 79, wherein the enzyme is glucose dehydrogenase.

81. The sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62, 66, 69-71, 76, 78 and 80, wherein the enzyme is equally distributed throughout the sensing layer (9).

82. A sensor (1) according to claim 81, wherein the sensing layer (9) has an effective diffusion coefficient D_effThe enzyme loading parameter α of 1/10-1/200.

83. Sensor (1) according to any one of claims 1-5, 7-9, 11, 13-15, 18-20, 22-24, 26-28, 30-32, 34-39, 41, 43, 45, 47, 49, 51, 54-58, 60-62, 66, 69-71, 76, 78, 80 and 82, wherein the sensing layer (9) is arranged on the contact pad (7) and the sensing layer (9) has a lower surface facing the contact pad (7) and an upper surface facing away from the contact pad (7).

84. A method of manufacturing an amperometric sensor (1) configured for implantation into the living body of a human or animal for measuring the concentration of an analyte in a body fluid, the method comprising the steps of:

-mixing carbon particles, an enzyme and a polymeric binder to form a paste;

-applying the paste onto the support member (5) near the contact pads (7),

-hardening the paste into a porous sensing layer (9).

85. The method according to claim 84, wherein the cross-linked enzyme is mixed with the carbon particles and the polymeric binder.

86. A method according to claim 84 or 85, wherein the sensing layer (9) is manufactured such that it has an effective diffusion coefficient D_effCharacterised by diffusion of the analyte in the sensing layer (9) and is 1/10-1/1000 of the diffusion coefficient D of the analyte in water.

87. A method according to claim 84 or 85, wherein the sensing layer (9) is manufactured such that it is electrically conductive.