JP2008541104A - Enzyme sensor having a hydrous spacer layer - Google Patents

Abstract

The present disclosure is an amperometric enzyme sensor comprising a hydrous spacer layer in contact with an electrode, eg, a porous polymeric matrix spacer layer selected from polyethylene terephthalate (PETP), polyvinyl chloride and polycarbonate. About the sensor. The sensor is useful for measuring the presence or amount of biological analytes such as glucose, lactate, creatine, creatinine and the like.
[Selection] Figure 3

Description

  The present invention relates to an amperometric enzyme sensor having a hydrous spacer layer in contact with an electrode.

  An enzyme sensor is a sensor that causes a chemical species (analyte) to be measured to undergo an enzyme-catalyzed reaction in the sensor before detection. Reaction of the analyte with the enzyme (if the analyte is a substrate) or enzyme cascade results in a second species whose concentration (under ideal conditions) is proportional to or consistent with the concentration of the analyte. Next, the concentration of the second type is detected by a converter, for example, by an electrode.

  Enzyme of an enzyme sensor is typically included in a sensor membrane suitable for contacting a fluid sample. Most typically, the enzyme is contained in a separate enzyme layer of the sensor membrane, which is separated from the fluid sample by a cover membrane. Thus, the analyte contacts the enzyme after diffusion through the sensor cover membrane, then the enzyme / analyte reaction occurs, and then the second species diffuses into the detector portion of the sensor, eg, the electrode. This produces a response related to the analyte concentration.

On the other hand, the enzyme layer is in most cases separated from the electrode by an interference limiting layer that allows a second type of diffusion. Conventional enzyme-based H ₂ O ₂ detection sensors typically include a cover membrane for the enzyme layer (ie, a diffusion limiting layer), an enzyme layer, an intermediate layer (ie, an interference limiting layer), and a metal anode. . Such a system works satisfactorily when used with analytes that are generally present at high concentrations (eg, lactate and glucose in a blood sample). However, if the system is used to detect analytes that are present at very low concentrations (eg, creatinine, creatine, or other analytes having a detection limit in the range of about 1 to about 20 μM), It has been observed that an analyte-free fluid sample can produce a significant spurious signal on the electrode. These pseudo-signals may correspond to signals of about −25 μM to about 25 μM analyte, which are various liquids brought into contact with the enzyme sensor, such as blood samples, washings, wettings, calibrations It originates in the compositional difference of the liquid etc. (other than the subject that is zero). Its operating state is probably a combination of two different causes.

First, non-ionic species diffuse more rapidly than ionic species across the interference limiting layer. Thus, bicarbonate / CO ₂ present in the sample but not in the rinse solution will lower the pH below the interference limiting layer. The same behavior is seen with imidazole / H-imidazole present in most rinse solutions but not in the sample. The drop in pH will cause a zero current drop resulting from the oxidation of water.

  Second, the concentration of ionic species at the anode surface varies as a function of different samples. Such a change will result in a change in the ionic composition on the electrode, thereby resulting in a current known as a non-Faraday current. The total amount of charge delivered as a non-Faraday current will only depend on the difference in ionic composition, however, the diffusion time constant may vary.

The pseudo signal is particularly problematic for differential measurement when measuring the concentration of an analyte such as creatinine in a blood sample.
Therefore, there is a need for an enzyme sensor that reduces or even eliminates the spurious signal that results from the use of liquids of different compositions.

  US 2004/0011671 A1 discloses a device and method for measuring analyte levels, specifically an implantable device for monitoring glucose levels in body fluids.

  WO 90/05910 A1 discloses a completely microfabricated biosensor that includes an analyte attenuation layer.

  It has been found that the above problem can be alleviated by introducing a hydrous spacer layer between the anode and the interference limiting layer. More specifically, the present invention relates to an enzyme sensor for measuring the concentration of an analyte in a fluid sample, comprising: an electrode; a hydrated spacer layer in contact with the electrode; at least one intermediate layer. The innermost part of the at least one intermediate layer is in contact with the spacer layer; and at least one enzyme layer, wherein the innermost part of the at least one enzyme layer is of the at least one intermediate layer It has been found that by providing an enzyme sensor that includes the most in contact with the outside, higher accuracy and reliability of the enzyme sensor can be achieved.

DETAILED DESCRIPTION OF THE INVENTION As described herein, the present invention relates to an amperometric enzyme sensor that measures the concentration of an analyte in a fluid sample.

  As described herein, the term “enzyme sensor” generally encompasses an electrochemical sensor that includes an enzyme (or enzyme cascade) that can convert the analyte of interest into a second species. It is. The analyte in question is a potential fluid sample component, i.e. an enzyme sensor is typically used to measure the concentration of the analyte in the fluid sample. “Analyte” is sometimes also referred to as “enzyme substrate” or simply “substrate”.

  The sensor of the present invention is typically a multi-use sensor. A multi-use sensor is to be understood as a sensor that is used for more than one measurement, and therefore may be exposed to more than one sample volume and in intermittent contact with a calibration solution, a washing solution, and the like. Such sensors are usually used for a longer time.

  The fluid sample can be primarily any liquid (preferably an aqueous liquid) that is compatible with the sensor, specifically the cover membrane. Fluid samples include body fluids such as urine, saliva, interstitial fluid, spinal fluid and blood. The blood includes whole blood samples, diluted blood samples, blood fractions, pre-reaction blood samples, and the like. The sensor is particularly well suited for whole blood samples.

  The sensor of the present invention may be of a conventional type or of a flat type, for example a thick film sensor or a thin film sensor. The enzyme membrane of such a sensor is often referred to as a laminated membrane.

  In the case of conventional types of enzyme sensors, membranes, such as multilayers, including, for example, spacer layers, intermediate layers, enzyme layers and cover membranes, are typically assembled after being assembled as discrete objects, It is placed in connection with the electrode (ie, generally attached to its tip). For example, see FIG. Such a method for constructing a multilayer film is well known in the art. See, for example, WO 98/21356. Conventional types of enzyme sensors can include track-etched membranes as well as solution cast membranes.

  In the case of flat type enzyme sensors, for example, thick film sensors and thin film sensors, the enzyme film and electrode including the water-containing spacer layer are made of materials corresponding to the electrode, spacer layer, intermediate layer, enzyme layer and cover layer, respectively (typically In a sequential and individual manner by depositing on a solid dielectric support, such as a ceramic or wafer material. An example of planar sensor construction is shown in FIG. Methods for constructing planar type sensors, such as thick film sensors and thin film sensors, are well known in the art. See, for example, WO 01/90733, WO 01/65247, and WO 90/05910. The material corresponding to such a sensor membrane layer is in most cases deposited by solution casting.

  The enzyme sensor of the present invention comprises an electrode; a hydrous spacer layer in contact with the electrode; at least one intermediate layer, the innermost part of the at least one intermediate layer being in contact with the spacer layer; And at least one enzyme layer, wherein the innermost part of the at least one enzyme layer is in contact with the outermost part of the at least one intermediate layer.

  The enzyme sensor spacer layer is generally ready for use, i.e. the hydrous spacer layer contains a substantial amount of water, in which case the enzyme sensor measures the analyte in the fluid sample. It is stated that it is possible. However, the enzyme sensor is stored in a dry form, i.e. with the spacer layer substantially dry and supplied to the end user. Thus, the end user will need to wet the membrane of the enzyme sensor with an aqueous liquid so that the spacer layer capable of absorbing a significant amount of water is converted to a hydrous porous spacer layer. Other layers may also be able to absorb a significant amount of water.

  In conventional enzyme sensor construction, wetting can typically be done with the internal fluid of the enzyme sensor (see, eg, FIG. 1). The flat sensor is typically wetted by, for example, a specific wetting liquid, a washing liquid or a calibration liquid.

  The main members of the enzyme sensor are (i) an electrode, (ii) a hydrous spacer layer in contact with the electrode, (iii) at least one intermediate layer, and the innermost part of the at least one intermediate layer is Is in contact with the spacer layer, and (iv) at least one enzyme layer, the innermost part of the at least one enzyme layer being in contact with the outermost part of the at least one intermediate layer . The spacer layer, at least one intermediate layer, at least one enzyme layer and optionally the cover layer together form an enzyme membrane of the enzyme sensor. These individual members will be described in detail below.

Water-containing spacer layer The enzyme sensor of the present invention includes a water-containing spacer layer that separates the electrode from at least one intermediate layer.

The term “hydrated spacer layer” refers to a layer that provides a buffering effect in reducing pH instability at the electrode surface when the sensor is used.
Since the water in the hydrous spacer layer buffers the change in ionic composition experienced by the anode, it is believed that the non-Faraday current extends over a longer time interval and results in a current having a smaller amplitude. Diffusion is a very rapid process over a small distance (typically less than about 1 s for diffusion of O ₂ over about 50 μM). Thus, the spacer layer does not function in isolation, but only in combination with a diffused resistor (eg, an interference limiting layer) so that the system functions like a series capacitor that includes a resistor. . As such, the interference limiting layer should be rather impermeable to ions or else the spacer layer should be very thick.

The high moisture content of the spacer layer ensures that the analyte (eg, H ₂ O ₂ ) can diffuse easily and rapidly across that layer. Computer modeling of sensor signals with and without a water-containing spacer layer supports these findings. For example, for the sensor system described below, an amplitude decrease of less than 5% and a time constant increase of 13.0s to 13.2s were observed.

  Suitable examples of materials that form a porous polymeric matrix of a hydrous spacer layer for conventional sensors (track etched or solution cast) include polyethylene terephthalate (PETP), glycol modified polyethylene terephthalate (PETG) and Polyesters such as glycol modified polycyclohexylenedimethylene terephthalate (PCTG), polycarbonate, cellulose (eg, regenerated, acetate, triacetate, acetate butyrate), polyolefins and their derivatives, fluorinated hydrocarbon polymers and copolymers (eg, polychloro Trifluoroethylene, polyvinylidene fluoride, polytetrafluoroethylene, polyethylene chlorotrifluoroethylene, polyethylene tetrafluoroethylene , Fluorinated ethylene propylene copolymer), polyimide (eg, Kapton), polystyrene, poly (meth) acrylate, polyvinyl chloride and its derivatives (including copolymers such as vinyl chloride co (meth) acrylate type copolymer), polyamide, polyurethane, Copolymers of polysulfone, polyethersulfone, polyphenylene sulfide, silicone, and organosiloxane polycarbonates (for example, those disclosed in US 3,189,662), particularly including polyethylene terephthalate (PET), polyvinyl chloride and polycarbonate. However, it is not limited to this. In one embodiment, the material comprising the space layer for such a sensor is polyethylene terephthalate (PETP). Preferably, such a spacer layer has been track etched.

Suitable examples of (solution casting) materials that form a porous polymeric matrix of a hydrous spacer layer for planar sensors, such as thick film sensors, include hydrophilic polyurethanes, hydrophilic poly (meth) acrylates, poly (vinyl) Pyrrolidone), polyurethanes, Nafion ^™ polymers, electropolymerized polymers (eg, polythiophene, 1,3-diaminobenzene, phenol) and polymers selected from SPEES-PES (polyarylethersulfone / polyethersulfone copolymers). However, it is not limited to this. Alternatively, the material forming the porous polymeric matrix is the same material as described immediately above for the conventional sensor spacer layer mixed with a porous forming compound (eg, a detergent or a water soluble hydrophilic polymer), Specifically, it can be selected from polyvinyl chloride and polycarbonate mixed with such a porosity forming compound.

  As used herein, the term “hydrated” as used with respect to the spacer layer means that the porous polymeric matrix contains a substantial amount of water, eg, an amount of at least 6% water based on the weight of the porous polymeric matrix. It means that. The moisture content may be even higher, for example at least 8%, such as at least 10% or at least 20%, or at least 25%, or at least 40% or even at least 50%, or higher. It may be. For solution cast planar sensors, the degree of total swelling (ie, water absorption) should be carefully considered because excessively high water absorption may be detrimental to the structural integrity of the enzyme membrane. Thus, for planar sensors, the moisture content should preferably not exceed 200%, such as 150%.

In order to further reduce the effect of bicarbonate (HCO ₃ ⁻ ) in fluid samples and other liquids, it may be advantageous to include a buffer, cation exchange material or electrolyte salt in the spacer layer. . Thus, in one embodiment, the hydrous spacer layer further comprises one or more components selected from buffering agents, electrolyte salts (eg, electrolyte polymers) and cation exchange materials.

  The spacer layer can have a porosity in the range of 0.0005-2% (vol / vol) for the track-etched material and in the range of 1-90% for the solution cast material.

For conventional creatinine / creatine and urea sensors that include a track-etched spacer layer, the porosity is preferably in the range of 0.05-0.1%, such as 0.2-0.25%. It may be. For conventional lactate sensors that include a track-etched spacer layer, the porosity may preferably be in the range of 0.0005 to 0.015%, such as 0.003 to 0.004%. For conventional glucose sensors that include a track-etched spacer layer, the porosity may preferably be in the range of 0.001-0.05%, such as 0.01-0.02%. The porosity of the track-etched film is determined as porosity (%) = π × (pore diameter / 2) ² × (pore density) × 100%. The average pore size of the spacer layer may be in the range of 0.05 to 250 nm, such as 1 to 150 nm or 10 to 110 nm, and the pore density is 40,000 to 40,000,000 fine. It may be in the range of holes / cm ² .

  The porosity of the solution cast spacer layer can be most easily determined as the volume occupied by water when the membrane is wet with water. The porosity of the solution cast spacer layer may be in the range of 1-90%, such as 3-85%. The at least one order of difference between the porosity of the track-etched membrane and that of the solution cast membrane is explained by considering only the “effective” pores of the track-etched membrane. Although determination of the porosity of a solution cast membrane can include all pores and cavities.

The weight ratio between the water and the solids in the spacer layer may be in the range of 10: 1 to 1:10, for example 8: 1 to 1: 5 or 10: 1 to 1: 2.
The hydrous spacer layer typically has a thickness in the range of 0.2-20 μm, such as 0.5-15 μm. For planar sensors, the thickness is typically in the range of 0.2-10 μm, such as 0.5-5 μm. For conventional sensors, the thickness is typically in the range of 1-20 μm, such as 2-15 μm.

  The hydrous spacer layer may be in the form of a solution cast layer or a track etched film. In order to obtain appropriate porosity, specifically for spacer layers for flat sensors, eg, thick film sensors, the polymeric materials described above and porosity forming compounds (eg, detergents, water soluble hydrophilics) It is often desirable to mix them with a functional polymer.

  In the case of conventional sensors, the pores are oriented substantially perpendicular to the electrode surface so that the second species detected at the electrode is more accurately directed to the electrode surface (eg, It is preferred to use a track-etched film since it is considered important (see FIG. 1). This causes reduced diffusion from the boundary area and thus gives a faster sensor response. Further, with this construction, radial electrical contact is minimized, thereby reducing external electrical noise.

  In one embodiment, the sensor is of a conventional type and the hydrous spacer layer is a track-etched polyethylene terephthalate material. In another embodiment, the sensor is of the planar type and the hydrous spacer layer is preferably a hydrophilic polyurethane or mixed with a porous forming compound selected from, for example, detergents, water soluble hydrophilic polymers, etc. It is a solution cast layer of hydrophilic poly (meth) acrylate.

  For the various embodiments defined above, the combination of at least one intermediate layer and a water-containing spacer layer that isolates the electrode from at least one enzyme layer allows diffusion of compounds such as paracetamol, ascorbic acid and uric acid in the first 15 seconds. It is preferred that the signal can be limited in such a way as to reduce the signal by at least 90%, such as at least 95%.

Intermediate layer In one embodiment, the enzyme layer is not in direct contact with the spacer layer, so the enzyme sensor may preferably include at least one intermediate layer that functions as an interference limiting layer.

In another embodiment, the at least one intermediate layer comprises, for example, cellulose acetate (CA), Nafion ^™ , rigid PVC, Baytron ^™ , electropolymerized polymer (eg, polythiophene, 1,3-diaminobenzene, phenol) and SPEES- Selected from PES (polyarylethersulfone / polyethersulfone copolymer). In one embodiment, at least one intermediate layer is an interference limited cellulose acetate (CA) layer.

Combination of spacer layer and intermediate layer In another variant, the hydrous spacer layer and at least one intermediate layer are combined into a material of the type described for the intermediate layer and a heterogeneous layer of the type of material described for the spacer layer. . The layer is formed in such a way that the spacer layer type material is dispersed in a continuous phase of at least one intermediate layer type material. Options for materials and properties are as described above for the spacer layer and at least one intermediate layer.

Electrode The electrode of the enzyme sensor is selected in consideration of the reaction product of the analyte and one or more enzymes (for example, an enzyme cascade as in the case of the creatinine sensor). Typically, the electrodes are made from noble metals such as gold, palladium, platinum, rhodium, indium or iridium, preferably gold or platinum, or mixtures thereof. Other suitable electron-conductive material, MnO _2, Prussian blue, graphite, iron, include nickel and stainless steel.

In some cases it is preferred to further include additional electrodes adjacent to the required electrode, eg, an internal reference electrode and / or a counter electrode. For example, see FIG.
Enzyme layer The enzyme layer (s) of the enzyme sensor facilitates the conversion of the analyte into a secondary species that one or more enzymes can detect on the electrode surface, Plays an important role. In some embodiments, a single enzyme is used (eg, glucose oxidase or lactate oxidase, urease), but multiple enzymes (eg, creatinase or sarcosine oxidase) are used to provide a reaction cascade that results in a species that can be detected. Can be made easier.

  One or more enzymes may be attached by themselves or fixed directly or indirectly, such as embedded or mixed in a polymer, or to reduce or eliminate migration It may be in a cross-linked or fixed form to the base layer or cover layer. In some embodiments, multiple enzymes can be placed in separate layers. The enzyme layer is further in place by a ring or gasket to avoid the use of excess amounts of enzyme and to ensure that a well-measured amount of enzyme is placed in a well-defined area of the sensor membrane. Can be held.

  At least one enzyme layer is composed of carbohydrate oxidase, glucose oxidase, galactose oxidase, glycolate oxidase, aldose oxidase, pyranose oxidase, lactate oxidase, α-hydroxy acid oxidase, sarcosine oxidase, alcohol oxidase, glycerol oxidase, amine oxidase, amino acid oxidase, Includes, but is not limited to, cholesterol oxidase, urease, bilirubin oxidase, laccase, peroxidase, glucose dehydrogenase, lactate dehydrogenase, glutamate dehydrogenase, P-450, superoxide dismutase, catalase, creatininase, creatinase and related coenzymes Not small Kutomo may include one enzyme.

  For detection of creatine, the enzyme layer preferably comprises creatinase and sarcosine oxidase. For detection of creatinine, the enzyme layer preferably comprises creatininase, creatinase and sarcosine oxidase. For the detection of glucose, the enzyme layer preferably comprises glucose oxidase. For detection of lactate, the enzyme layer preferably comprises lactate oxidase. For the detection of urea, the enzyme layer preferably contains urease.

Cover membrane The enzyme sensor further provides at least one porous polymeric material to ensure that a certain well-defined typical amount of analyte diffuses into the enzyme layer, ie under controlled conditions. Cover membranes may be included. Such analyte-limited conversion is a prerequisite for obtaining a substantially linear relationship between sensor response and analyte concentration within a reasonable range.

Thus, the enzyme sensor of the present invention preferably includes a diffusion limiting layer in the form of a cover membrane adapted to separate the enzyme layer from the fluid sample. The cover membrane is preferably such that it does not exceed the capacity of the immobilized enzyme for analyte conversion and that there is sufficient oxygen (O ₂ ) in the enzyme layer for enzyme conversion of the analyte. It is a porous membrane that limits the diffusion of the analyte into the enzyme layer. The principle of the diffusion limiting layer is described, for example, in Danish Patent No. 170103. Thus, in fact, any known cover membrane can be useful with the enzyme sensor of the present invention.

  Furthermore, analyte diffusion through the cover membrane is invariant over time and from sample to sample so that the same analyte concentration in two separate fluid samples results in a well-defined sensor response. Is desired. What is further desired is that the cover membrane is capable of rapidly diffusing a small amount of analyte across the membrane, thereby facilitating uniform dispersion of the analyte in the enzyme layer. Thus, the enzyme in the enzyme layer quickly converts the analyte to the second species, producing a rapid sensor response. Such nearly simultaneous conversion of the analyte results in improved linearity. Furthermore, it is desirable that macromolecules such as proteins and enzymes are substantially prevented from moving across the cover membrane. For example, proteases present in a wash solution or blood sample will have a detrimental effect on the enzyme layer if such proteases are able to migrate into and through the cover membrane. It has been noticed.

On the other hand, it is also important that the cover membrane can provide high retention of the second species (eg, H ₂ O ₂ and O ₂ ) in the sensor, so that the second species The response derived from is not biased by a significant amount of those species diffusing through the pores of the cover membrane, and is sufficient to maintain a conversion limited by the analyte. O ₂ is retained in the enzyme layer. These characteristics are particularly relevant in considering whether an intermediate layer having a diffusion limiting effect is arranged between the enzyme layer and the electrode.

  The at least one porous polymeric material can be selected from a fairly wide range of materials. Representative examples include polyesters such as polyethylene terephthalate (PETP), glycol modified polyethylene terephthalate (PETG) and glycol modified polycyclohexylenedimethylene terephthalate (PCTG), polycarbonate, cellulose (eg, recycled, acetate, triacetate, acetate butyrate). Rate), polyolefins and their derivatives, fluorinated hydrocarbon polymers and copolymers (eg, polychlorotrifluoroethylene, polyvinylidene fluoride, polytetrafluoroethylene, polyethylene chlorotrifluoroethylene, polyethylene tetrafluoroethylene, fluorinated ethylene propylene copolymers) ), Polyimide (eg, Kapton), polystyrene, poly (meth) acrylate, polyvinyl chloride And derivatives thereof (including copolymers such as vinyl chloride co (meth) acrylate type copolymers), copolymers of polyamide, polyurethane, polysulfone, polyethersulfone, polyphenylene sulfide, silicone, and organosiloxane polycarbonate (eg, US Pat. No. 3,189,662). Are disclosed).

In one aspect of the invention, the at least one porous polymeric material is selected from polyethylene terephthalate (PETP), polyvinyl chloride and polycarbonate.
In one embodiment, the porous polymeric material is polyethylene terephthalate (PETP).

In another aspect, the porous polymeric material is polyvinyl chloride (PVC).
Typically, the at least one porous polymeric material does not comprise a hydrophilic polyurethane, which is an excess level of H, particularly if such a material includes an interlayer that has a diffusion limiting effect. _This is because it is considered that ₂ O ₂ diffusion will be given.

  In one embodiment, a cover membrane having some of these advantageous properties has an outer surface and pore openings on at least one face of at least one porous polymeric material, preferably hydrophilic polyurethane and hydrophilic poly This is achieved by coating with a hydrophilic polymer selected from (meth) acrylates.

  The cover membrane (and also the at least one porous polymeric material) includes two sides: one side proximate to the enzyme layer and one side distal to the enzyme layer, the latter further comprising an enzyme sensor When in use, it faces the fluid sample. As described above, at least one surface of at least one porous polymeric material is coated with a hydrophilic polymer.

  In one aspect of the invention, at least the surface distal to the enzyme layer is coated with a hydrophilic polymer. This aspect provides benefits with respect to reducing or eliminating blood bias and blood drift and extends the life of the sensor.

  In another aspect, at least the surface proximate to the enzyme layer is coated with a hydrophilic polymer. It is believed that problems with changing sensitivity, lack of linearity, analyte distribution in the enzyme layer, and reduced lifetime can be reduced or eliminated in this manner. If at least one porous polymeric material is properly selected, e.g., by selecting a porous polymeric material that itself has the appropriate blood compatibility, problems with blood bias and blood drift can cause problems with the enzyme layer. Even in the absence of a hydrophilic polymer covering the outer surface and pore opening of the distal at least one porous polymeric material face, it can be reduced at least in part.

  In a preferred embodiment, both surfaces are coated with a hydrophilic polymer. This arrangement provides the benefit of improving analyte distribution, sensitivity and linearity in the enzyme layer, even reducing or eliminating blood bias and blood drift, extending the life of the sensor, limiting enzyme migration, and Improve linearity.

The expression “external surface and pore opening” refers to each of two faces of at least one porous polymeric material that is a surface interrupted by a pore opening (pore opening).
In the text, the expression “coated with” means that not only the surface of at least one porous polymeric material but also the pore opening is more hydrophilic than hydrophilic polymers (eg hydrophilic polyurethanes and hydrophilic poly (meth) acrylates). Selected).

  As used herein, the expression “hydrophilic polymer” is intended to mean a single hydrophilic polymer, as well as a mixture of two or more hydrophilic polymers. It should be understood that the one or more hydrophilic polymers described above can be mixed with up to 30% of other non-hydrophilic polymers. However, in one preferred embodiment, the coating on the at least one porous polymeric material only comprises a hydrophilic polymer selected from hydrophilic polyurethanes and hydrophilic poly (meth) acrylates.

  The use of a hydrophilic polymer, for example selected from hydrophilic polyurethanes and hydrophilic poly (meth) acrylates, to coat the outer surface and pore opening of at least one porous polymeric material is further prior to coating. The diffusivity can be adjusted to obtain the desired diffusion limit to produce enzyme membranes that depend on the different porosity of different porous polymeric materials and are suitable for different sample analyte concentration ranges.

  For planar sensors, a coating of a hydrophilic polymer (eg, selected from hydrophilic polyurethanes and hydrophilic poly (meth) acrylates) typically comprises a solution of the hydrophilic polymer with at least one porous polymeric material. It is obtained by distributing, spraying, screen printing, etc. on the surface (and pore opening). In one aspect of the invention, the outer surface and pore opening of at least one surface of at least one porous polymeric material distal to the enzyme layer is coated with a hydrophilic polymer. Another embodiment, ie, after the enzyme layer is coated with a hydrophilic polymer, at least one porous polymeric material is provided (and optionally the porous polymer material is coated with the same or another hydrophilic polymer). However, it is thought that it is exactly as the aspect by which both surfaces were coated.

  For conventional sensors, the coating on the at least one porous polymeric material allows a solution of the hydrophilic polymer to a surface (and only one or both sides) of the at least one porous polymeric material. It can be obtained by dispensing, spraying, screen printing, etc. on the pore openings) or at least one porous polymeric material can be immersed in a solution such as a hydrophilic polymer.

  Thus, specifically, for a conventional sensor comprising a track-etched porous material, both sides of at least one porous polymeric material may be coated with a hydrophilic polymer. The surface of the at least one porous polymeric material proximate to the enzyme layer may also be coated with a hydrophilic polymer, which means that a relatively large distance between every single pore means that there is no hydrophilic polymer coating. In particular, it is believed to provide special advantages for track-etched porous polymeric materials because they produce a non-linear response in the presence. In this situation, the analyte (enzyme substrate) should diffuse into unoccupied enzyme molecules in the enzyme layer, but a longer diffusion distance causes non-simultaneous conversion. In contrast, a hydrophilic polymer coating on the surface of at least one porous polymeric material proximate to the enzyme layer facilitates diffusion of the analyte within that layer and provides the analyte to the enzyme layer more uniformly. There will be a higher or more linear response. Thus, with a hydrophilic polymer coating on the surface of at least one porous polymer material proximate the enzyme layer, the analyte should travel a short distance in the denser enzyme layer until it reaches the available enzyme molecules. It will only be.

  In some aspects of the invention, not only the outer surface and pore openings of at least one porous polymeric material are coated with a hydrophilic polymer, but the hydrophilic polymer also has its at least one At least partially penetrates the pores of the porous polymeric material from the surface.

In these specific embodiments, at least one porous polymeric material of the cover membrane layer is said to be at least partially impregnated with a hydrophilic polymer.
As used herein, the term “impregnated” means that the hydrophilic polymer coats the outer surface and pore openings on both sides of at least one porous polymeric material, and further the pores of the porous polymeric material. It means that it is penetrating.

  The term “at least partially impregnated” means that the hydrophilic polymer covers the outer surface and the pore opening of at least one surface of at least one porous polymeric material, and further on at least one surface thereof. It means that at least a part of the pores of the resulting porous polymeric material has penetrated.

  In another aspect of the invention, the hydrophilic polymer is substantially insoluble in water when the sensor is used. However, the hydrophilic polymer is preferably not cross-linked when applied to and covers the cover membrane, and preferably no subsequent cross-linking occurs. Instead, the hydrophilicity and water insolubility of the hydrophilic polymer is obtained by a suitable combination of hydrophilic segment and hydrophobic segment / portion of the hydrophilic polymer. This arrangement provides a much simplified manufacturing procedure because the cross-linking step can be omitted entirely.

The term “insoluble in water” means a polymer that does not substantially dissolve in water when a cover membrane coated with a hydrophilic polymer is stored in an aqueous solution at 25 ° C. for 24 hours.
In certain embodiments of the invention, the at least one porous polymeric material has a porosity in the range of 0.002-30% (vol / vol).

  The desired porosity of the polymeric material depends in part on the desired upper limit of the detection range. The very high upper limit of the detection range will require a fairly low porosity in order to obtain a wide linear range where the cover membrane should give the analyte a rather high diffusion resistance. When expressed as the number product of porosity (% (vol / vol)) and upper limit of linear detection range (mM of analyte), the value is preferably 0.05 to 10 [% (vol / vol) .MM] or preferably within a range of 0.01 to 50 [% (vol / vol) .mM], such as 0.1 to 2 [% (vol / vol) .mM].

The average pore size of the porous polymeric material may be in the range of 0.05 to 250 nm, such as 1 to 150 nm or 10 to 110 nm.
In one embodiment, specifically when the sensor is of a conventional type, the porous polymeric material has a pore density in the range of 40,000 to 40,000,000 pores / cm ^2. A track-etched material having

For creatinine / creatine sensors and urea sensors that include track-etched cover membranes, the porosity is preferably in the range of 0.05-0.1%, such as 0.2-0.25%. For lactate sensors that include a track-etched cover film, the porosity is preferably in the range of 0.0005 to 0.015%, such as 0.003 to 0.004%. For a glucose sensor that includes a track-etched cover membrane, the porosity is preferably in the range of 0.001-0.05%, such as 0.01-0.02%. The porosity of the track-etched film is determined as porosity (%) = π × (pore diameter / 2) ² × (pore density) × 100%.

  The porosity of a solution cast membrane can be more easily determined as the volume occupied by water when the membrane is wet with water. The porosity of the solution cast membrane is typically in the range of 1-40%, such as 3-30%. At least one order of difference between the porosity of the track-etched membrane and the porosity of the solution cast membrane is only the “effective” pores (ie, the through-holes) of the track-etched membrane. However, all pores and cavities are included in the determination of the porosity of a solution cast membrane.

In one embodiment, the hydrophilic polymer is a hydrophilic polyurethane.
Polyurethane is the most widely used biomedical polymer for blood contact surfaces, for example, implants and medical devices. Polyurethane elastomers are multiphase block copolymers consisting of alternating blocks of hard and soft segments. The hydrophobic hard segment is formed by the reaction of an aliphatic, cycloaliphatic or aromatic diisocyanate with a diol, diamine or water. The soft hydrophilic or relatively hydrophilic segment is composed of low molecular weight hydroxy terminated polyethers, polyesters or aliphatic polyolefins. Hydrophilic polyols can be used as chain extenders or alternatively can be included in the prepolymer. Chemical incompatibility between the hard and soft segments and between the hydrophobic and hydrophilic segments leads to phase segregation in the polyurethane. The hard segment domains are interconnected by secondary bonds and dispersed in the soft segment matrix, but serve as physical bridges that reinforce the overall system. The soft matrix can be adjusted for hydrophilicity by using a mixture of polyethers or polyesters having different hydrophilicities. For very hydrophilic polyurethanes, polyethylene glycols are often used, but their hydrophilicity control can be achieved with higher polyalkyl ethers such as polypropylene glycol and polybutylene glycol. In this way, polyurethanes can be made to be hydrophilic, hydrophobic, hydrophilic / hydrophobic, hard and rigid or soft and elastic, hydrolytically stable, or deliberately degradable. Because of their hard and soft segment structure, these polyurethanes are mechanically strong, tear resistant and exhibit good flex life. These properties make them suitable as hydrophilic coatings for sensor membranes. These coatings have high water absorption due to the content of hydrophilic segments and good use stability. Because of their pseudo-crosslinked segment structure, they are also insoluble in water.

  The hydrophilic polyurethane can be selected from polyurethanes having hydrophilic segments included therein, for example, segments such as polyethylene glycol and polypropylene oxide. Such hydrophilic polyurethanes can be produced from polyalkylene glycols (polyalkylene oxides) that have terminal hydroxy groups or amino groups to form linear polymer chains by reaction with diisocyanates. Examples of such hydrophilic polyurethanes are those disclosed in US 4,789,720; US 4,798,876; and US 5,563,233. Other suitable examples include polyurethanes modified with hydrophilic groups such as aliphatic polyethers. See for example US 6,200,772 B1.

  The hydrophilic segment is typically derived from polyethylene glycol, amino-terminated polyethylene glycol, polypropylene glycol, amino-terminated polypropylene glycol, polyethylene oxide, polypropylene oxide and polyethyleneimine, especially polyethylene glycol.

  In some embodiments, the hydrophilic polyurethane is an aliphatic polyether urethane, an aliphatic polyether urethane urea, a cycloaliphatic polyether urethane, a cycloaliphatic polyether urethane urea, an aromatic polyether urethane, an aromatic polyether urethane. It is selected from urea, aliphatic polyester urethane, aliphatic polyester urethane urea, cycloaliphatic polyester urethane, cycloaliphatic polyester urethane urea, aromatic polyester urethane and aromatic polyester urethane urea. Aliphatic polyether urethanes or cycloaliphatic polyether urethanes (eg, cyclohexyl polyether urethane) are suitable as film coatings when linear or cyclic aliphatic diisocyanates are used. Naturally occurring isocyanates (eg, lysine diisocyanate) are also suitable. Cyclohexyl polyether urethane is believed to provide good biocompatibility to the membrane and to reduce or even eliminate membrane contamination.

In an aspect of the invention, the hydrophilic polyurethane backbone segments of polyethylene glycol _{- (CH 2 -CH 2 -O-)} n - and, specifically, at least 7% (w / w) or at least 10% (W / w) as a weight ratio of at least 5% (w / w) polyethylene glycol segments. It is believed that the inclusion of a substantial amount of polyethylene glycol segments imparts appropriate hydrophilicity and improves blood compatibility.

Suitable examples of preferred hydrophilic polyurethanes are those disclosed in US Pat. No. 5,322,063, incorporated herein in its entirety.
In one aspect of the invention, the hydrophilic polyurethane is a polysaccharide (eg, alginate, carrageenan, pectin and dextran), poly (HEMA), partially hydrolyzed polyvinyl acetate (PVA) or a cellulose derivative (eg, hydroxyethylmethylcellulose and carboxymethylcellulose). ) Main chain segments are typically at least 5% (w / w) polysaccharide, polyvinyl acetate or cellulose derivative segments, such as at least 7% (w / w) or at least 10% (w / w) Includes by weight.

  Examples of suitable commercially available hydrophilic polyurethanes include Hydromed D4 (water content when wet: 50% (w / w)) and Hydromed D640 (water content when wet: 93% (w / w)) Is included. Both polyurethanes are trade names of Cardiotech International Inc., Wilminton, MA, USA.

  Hydromed D4 and D640 products contain a central polybutylene oxide segment and polyalkylene oxide end groups. The polyalkylene oxide group may be polyethylene oxide or polyethylene oxide-polypropylene oxide-polyethylene oxide. In either case, the polyethylene oxide segment is preferably longer than the segment length of polybutylene oxide and polypropylene oxide. This is believed to provide sufficient hydrophilicity and water absorption as well as adequate blood compatibility.

In another embodiment, the hydrophilic polymer is a hydrophilic poly (meth) acrylate.
Examples of hydrophilic poly (meth) acrylates include a first monomer unit comprising an acrylate ester having a poly (ethylene oxide) substituent as part of the alcohol portion of the ester, and one or more selected from methacrylates and acrylates Acrylic copolymers containing a second monomer unit of The poly (ethylene oxide) substituent of the first monomer unit typically has an average molecular weight of 200 to 2000, such as 500 to 1500. Examples of such first monomers are methoxy poly (ethylene oxide) methacrylate, methoxy poly (ethylene oxide) acrylate, and the like. Examples of the second monomer unit include methyl methacrylate, ethyl acrylate, butyl methacrylate and the like.

Preferred hydrophilic poly (meth) acrylates include the acrylic copolymers disclosed in WO 93/15651 A1, which are incorporated herein in their entirety.
A preferred combination of monomers is methoxy poly (ethylene oxide) methacrylate, ethyl acrylate and methyl methacrylate.

Other preferred hydrophilic poly (meth) acrylates include those having poly (vinyl pyrrolidone) (PVP) segments or side chains.
In one aspect of the invention, the hydrophilicity of the hydrophilic polymer is such that the moisture content when wet is 5-100% (such as 25-95% (w / w) or 45-95% (w / w)). w / w) or 10-95% (w / w). Its water content is typically a function of the type and content of the hydrophilic polymer in the sense that higher hydrophilic segment content results in higher water content (when wet). Porosity can also play a role with respect to the preferred range of moisture content, ie, for membranes with small pores, the preferred range of moisture content is such as 10-30% (w / w). 5 to 80% (w / w) or 8 to 40% (w / w).

The nature of the cover membrane with respect to diffusion, diffusion rate, and in particular the ability to exclude large molecules is important for sensor functionality.
Also to be considered is the ability of the cover membrane to allow diffusion of glucose while limiting the diffusion of H ₂ O _{2 on} the other hand. Thus, in one embodiment, the diffusion rate of H ₂ O ₂ through the cover membrane relative to the diffusion rate of glucose through the cover membrane is in the range of 3-20, such as 3-15 or 3-10. . The diffusion rate is determined as described in the “Experiment” section. The relative diffusion rate of the cover membrane is superior to that of typical known polyurethane cover membranes.

The apparent diffusion coefficient of glucose through the cover membrane is preferably 0.1 to 1.5 × 10 ⁻⁹ or 0.5 to 1.1 × 10 ⁻⁹ , such as 0.3 to 1.5 × 10 ⁻⁹ for the glucose sensor. Should be in the range of 5.0 × 10 ⁻⁹ . For the corresponding lactate sensor, the apparent diffusion coefficient of lactate through the cover membrane is preferably in the range of 0.5-5 × 10 ⁻¹⁰ , such as 1.2-3.2 × 10 ^−10. Should. The apparent diffusion coefficient is measured as described in the “Experiment” section.

  Further relevant is the exclusion of “large” molecules (eg peptides, proteins, and enzymes such as enzymes in the enzyme layer (eg glucose oxidase and lactate oxidase)), but at the same time the relevant analyte, For example, the ability of the cover membrane to allow diffusion of lactate, glucose, creatine, creatinine and the like. Such analytes typically have a molecular weight of up to about 200, while peptides, proteins and enzymes have molecular weights of about 300 for small peptides to thousands or more for proteins, for example, It can have a molecular weight of about 30,000 in the case of glucose oxidase. The cover membrane layer (when in wet form) typically has a thickness in the range of 5-40 μm, such as 6-30 μm or 10-17 μm, for conventional sensors. For thick film sensors, the (wet) cover film layer typically has a thickness in the range of 1-20 μm, such as 2-10 μm or 3-5 μm.

The hydrophilic polymer layer of the dry cover membrane has a thickness in the range of 0.1-5 μm, such as 0.25-3 μm or 0.5-1 μm, especially for thick film sensors.
That is, the hydrophilic polymer layer of the dry cover membrane typically has a thickness in the range of 100-2000%, such as 100-1000% or 200-500% of the average pore size of the polymeric material. Have

In view of the preferred water absorption, the ratio between the thickness of the wet cover membrane and the thickness of the dry cover membrane is preferably in the range of 2: 1 to 1: 1.
The ratio between the thickness of the hydrophilic polymer layer of the wet cover membrane and the thickness of the hydrophilic polymer layer of the dry cover membrane is in the range of 100: 1 to 1: 1 or 80: 1 to 2: 1. Is within.

  In some embodiments, particularly for conventional sensors, the ratio between the hydrophilic polymer layer thickness of the wet cover membrane and the hydrophilic polymer layer thickness of the dry cover membrane is 10: 1. It may be in the range of 20: 1 to 1.5: 1, such as in the range of ˜2: 1. In some other embodiments, particularly for conventional sensors including, for example, track-etched membranes, the thickness of the hydrophilic polymer layer of the wet cover membrane and the hydrophilic polymer layer of the dry cover membrane The ratio between thickness is preferably in the range of 80: 1 to 10: 1, such as in the range of 50: 1 to 30: 1.

  For example, for a flat sensor including a solution cast membrane, the ratio between the thickness of the hydrophilic polymer layer of the wet cover membrane and the thickness of the hydrophilic polymer layer of the dry cover membrane is preferably 6: Within the range of 10: 1 to 2: 1, such as within the range of 1-3: 1.

  In other embodiments, the weight ratio between the porous polymeric material and the hydrophilic polymer (non-wet state) is 100: 1 to 1: 1, such as 80: 1 to 10: 1 or 50: 1 to 30. : 1.

In one embodiment of the invention, the cover membrane is the outermost layer of the enzyme sensor.
Some advantages have been identified by coating the outer surface (and pore opening) of at least one face of at least one porous polymeric material, for example with hydrophilic polyurethane. For one thing, polyurethane inherently has small pores that effectively block the pores of porous polymeric materials for enzyme / protein penetration / migration, but less hydrophilic and diffusion of hydrophobic molecules Still possible. In addition, hydrophilic polyurethanes are usually insoluble in water, but the polyurethanes are swellable and can retain a significant amount of water. As such, leaching and modification of the polyurethane coating will be substantially absent during the lifetime of the sensor. The same applies for example to hydrophilic poly (meth) acrylates.

Preferred Embodiments One embodiment is an amperometric enzyme sensor that measures the concentration of creatine in a fluid sample comprising a metal electrode (eg, a platinum electrode); in contact with the metal electrode, such as polyethylene terephthalate (PETP). ), In particular a water-containing spacer layer of a track-etched PETP material; an interference control layer in contact with the spacer layer (eg an interference limiting layer of cellulose acetate (CA)); an enzyme in contact with the cellulose acetate layer A layer (eg, an enzyme layer comprising sarcosine oxidase and creatinase); and a cover membrane layer for the enzyme layer, wherein the cover membrane layer comprises a porous polyethylene terephthalate material and wherein at least one At least of two porous polymeric materials The outer surface and the pore opening of the other surface are coated with a hydrophilic polyurethane comprising a polyethylene glycol main chain segment in a weight ratio of at least 5% (w / w) polyethylene glycol segments, and / or It relates to a sensor having a wet moisture content of at least 25% (w / w).

  Another aspect is an amperometric enzyme sensor that measures the concentration of creatinine in a fluid sample, comprising a metal electrode (eg, a platinum electrode); a hydrous spacer layer in contact with the metal electrode (eg, polyethylene terephthalate ( PETP), in particular a water-containing spacer layer of a track-etched PETP material); an interference control layer in contact with the spacer layer (eg an interference limiting layer of cellulose acetate (CA)); in contact with the cellulose acetate layer An enzyme layer (eg, an enzyme layer comprising sarcosine oxidase, creatininase and creatinase); and a cover membrane layer for the enzyme layer, wherein the cover membrane layer comprises a porous polyethylene terephthalate material And here, at least one porous The outer surface and pore opening of at least one side of the polymeric material is coated with a hydrophilic polyurethane comprising a polyethylene glycol main chain segment in a weight ratio of at least 5% (w / w) polyethylene glycol segments; And / or relates to a sensor having a wet moisture content of at least 25% (w / w).

Use of Enzyme Sensor The enzyme sensor of the present invention may be exposed to a wetting or calibration solution, usually until its signal is stable, before its first use.

  Measurement of, for example, creatinine, creatine, glucose, lactate, etc. in a body fluid sample can be performed with various automatic or semi-automated analyzers, many of which measure multiple parameters using multiple sensors. One example is a clinical analyzer, in particular a blood analyzer. The fluid sample is introduced manually or automatically into the flow system of the analyzer or into the flow system of the cassette for introduction into the analyzer. Accordingly, a sensor of one or more physiological sample parameters can be exposed to a fluid sample introduced into the flow system.

Accordingly, the present invention further provides an apparatus for measuring the concentration of an analyte in a fluid sample, the apparatus comprising one or more enzyme sensors as described herein.
Furthermore, the present invention is a method for measuring the concentration of an analyte in a fluid sample, comprising contacting the fluid sample with an enzyme sensor as described herein and requiring at least an electrode of the enzyme sensor. Further provided is a method comprising the step of taking a single measurement.

These sensors are typically exposed to samples directed to the sensor and samples derived from the sensor and other fluids such as wetting liquids, washing liquids, calibration liquids, and the like.
The above description is not intended in any way to limit the claimed invention. Furthermore, the combination of features discussed is not absolutely necessary for the interpretation of the present invention. Further, the disclosures of all patents or published applications cited herein are hereby incorporated by reference in their entirety.

  The invention is illustrated in more detail in the following examples. However, it should be understood that these examples are for illustrative purposes only and should not be used in any way to limit the scope of the present invention.

Experimental Materials Creatininase from Pseudomonas putida was obtained from Roche Diagnostics, Mannheim, Germany. Hydromed D4, Hydromed D640 and Hydromed TP were obtained from Cardiotech International Inc., Wilminton, MA, USA.

General procedure
The measured diffusivity of the apparent diffusion coefficient can be determined in the diffusion cell, in which case the value of the apparent diffusion coefficient of the substrate is obtained as a result of the total porosity and the diffusion coefficient of the substrate in water. “Apparent diffusion coefficient” means an “effective” diffusion coefficient of the entire membrane area that does not take into account the porosity of the membrane.

The diffusion cell (15 mm diameter, with O-ring) should be completely clean before use. In order to reduce contamination, it is desirable to have a high substrate concentration solution in the cell half where the O-ring is located. A fluid sample for analysis is filled in half of the cell without the O-ring. A film sample approximately ½ cm larger than the opening between the halves of the cells is cut and placed on the O-ring. The cell is then closed and sealed. A diffusion cell containing a membrane and a magnetic bar (10 mm) is placed on a magnetic stirrer (320 ± 30 rpm). The substrate solution in 30 mL of cleaning solution (S4970) and 30 mL of pure cleaning solution (S4970) are simultaneously filled in two half cells of the diffusion cell. After 48 hours and 72 hours, respectively, 1 mL of the sample is taken out with a syringe, and the substrate concentration in the pure cleaning solution is measured by filling with 1 mL of the pure cleaning solution. The substrate concentration in the sample is measured with an ABL ^™ 735 Blood Gas Analyzer (Radiometer Medical ApS, Copenhagen, Denmark).

The apparent diffusion coefficient is determined as follows. Flux: In all systems, the passive transport process of a compound will occur when the distribution of the compound in the system does not correspond to the thermodynamic equilibrium distribution of the compound. The flux is defined as the amount of compound that passes through area units (area perpendicular to the transport direction) per second and has units of J = quantity · cm ⁻² · s ⁻¹ .

  Fick's first law applies to stationary diffusion. That is, a linear concentration gradient was established.

Where D is the diffusion coefficient of the compound, ie, a value specific to the diffusible molecular type under the given conditions (if it only includes factors that determine the transport rate, such as size and shape) DC / dx is the gradient of the concentration profile at point x (dC / dx value is also called the concentration gradient in the x direction, in which case The sign letter indicates the direction in which the density increases, ie a positive value of dC / dx indicates that the density increases in the positive direction of the x-axis).

General sensor construction (conventional sensor type)
With reference to FIG. 1, the sensor 1 includes an electrode 2 on which a membrane ring 3 is attached. The electrode 2 includes a platinum anode 4 connected to a platinum wire 5 connected to a silver anode contact 7 via a microplug 6. The platinum anode 4 and the lower part of the platinum wire are sealed in a glass body 8. Between the glass body 8 and the micro plug 6, the platinum wire 5 is protected by a heat shrinkable tube. The tubular silver reference electrode 10 extends the length of the electrode 2 to the anode contact 7 which surrounds the upper part of the glass body 8 and is fixed inside the reference electrode by a fixing body 11 and an epoxy 12. The lower part of the glass body 8 is surrounded by the electrode base material 13, and the membrane ring 3 is attached to it.

  With reference to FIGS. 1 and 2, the upper part of the reference electrode 10 is used by the plug member 14 to mount the electrode 2 in the corresponding plug of the analyzer (not shown) and to fix the mantle 15 Surrounded. Gaskets 16 and 17 are placed between electrode 2 and mantle 15 to ensure that any electrolyte on the measuring surface of electrode 2 does not evaporate. The membrane ring 3 is attached to one end of the mantle 15 and includes a ring 20. The membrane 21 is stretched over the lower opening of the ring 20. This membrane 21 is shown in detail in FIG. 2, but as described in detail in Example 1.

General sensor construction (thick film sensor type)
FIG. 3 shows an exemplary planar thick film sensor construction formed on a dielectric support (110) on which a working electrode (120) and a reference electrode (130; 140) are formed. The electrodes are bounded by a two-layer dielectric encapsulant (150; 160 and 151; 161). The working electrode is covered with a water-containing spacer layer (121), an intermediate layer (170), an enzyme layer (180) and a cover membrane (190) disclosed herein.

  FIG. 3 shows an exemplary planar thick film sensor construction formed on a dielectric support (110) on which a working electrode (120) and a reference electrode (130; 140) are formed. The electrodes are bounded by a two-layer dielectric encapsulant (150; 160 and 151; 161). The working electrode is coated with a hydrated porous spacer layer (121), an intermediate layer (170), an enzyme layer (180) and a cover membrane (190) as disclosed herein.

  Referring to FIG. 3, one surface of a 200 μm thick alumina support 110 is coated with a circular platinum working electrode 120 having a diameter of 1000 μm and a thickness of 10 μm; an angular range of 30 to 330 ° of the outer periphery of the working electrode. A circular platinum counter electrode 130 with an outer diameter of 3000 μm, an inner diameter of 2000 μm and a thickness of 10 μm; and a circular silver / silver chloride reference electrode 140 with a diameter of 50 μm, located at 0 ° of the outer periphery of the working electrode. All three of these electrode structures are connected across the alumina support 110 to the sensor electronics (not shown) through a platinum-finished (filed) through hole (not shown) across the support. Yes. In operation, the working electrode 120 is polarized to +675 mV with respect to the reference electrode 140.

  Furthermore, on the alumina support 110, there is a two-layer structure of glass and polymer encapsulant. These two-layer structures comprise an outer diameter 1800 μm, an inner diameter 1200 μm and a 50 μm thick annular structure 160, 161 surrounding the working electrode 120; and a 50 μm thick structure 150, 151 surrounding the complete electrode system. Include. Both of these two-layer structures face the alumina support 110 of ESL glass 4904 from ESL Europe of United Kingdom, 20 μm thick inner layers 150, 160; and an international patent of SenDx Medical Inc. of California, USA 28.1 wt% polyethyl methacrylate (Elvacite, product number 2041, manufactured by DuPont), 36.4 wt% carbitol acetate, 34.3 wt% silanization as disclosed in application WO 97/43634 Outer layers 151, 161 of polymer encapsulant from SenDx Medical Inc. of California, USA containing kaolin (product number HF900 from Engelhard), 0.2 wt% fumed silica and 1.0 wt% trimethoxysilane. Consists of.

  The hydrous porous spacer layer 121 was formed by dispensing 300 nL of a 7% D4 PUR (Hydromed inc.) Solution in 96% ethanol onto the Pt working electrode by microdispensing.

  A circular inner membrane 170 of cellulose acetate and cellulose acetate butyrate having a diameter of 1200 μm and a thickness of 1 μm covers the working electrode 120 manufactured on top of the spacer layer 121. It is important that the membrane 170 covers all of the spacer layer, otherwise the interference removal will be incomplete.

A circular enzyme layer 180 of glucose oxidase crosslinked with glutaraldehyde having a thickness of 1200 μm and 2 μm in diameter covers the inner membrane 170.
The enzyme layer 180 was produced by dispensing 0.4 μl of a buffered solution of glucose oxidase crosslinked with glutaraldehyde onto the cellulose acetate membrane 170. The enzyme layer was dried at 37 ° C. for 30 minutes.

  A circular cover membrane layer 190 of PVC / trimethylnonyltriethylene glycol / diethylene glycol, having a diameter of 4000 μm and a thickness of 10 μm, covers the complete electrode system centered on the working electrode 120.

  The cover membranes consisted of 1.35 grams of polyvinyl chloride (Aldrich 34, 676-4), 0.0149 grams of trimethylnonyltriethylene glycol (Tergitol TMN3 from Th. Goldschmidt) and 0.134 grams of diethylene glycol. In addition to 21.3 grams of tetrahydrofuran and 7.58 grams of cyclohexanone. The mixture was stirred until the PVC dissolved and a homogeneous solution was obtained. 28.5 grams of tetrahydrofuran was added to give a 2% solution of 90/1/9 PVC / surfactant / hydrophilic compound composition. The solution was dispensed over the sensor area to cover all three types of electrodes and to have an approximately 0.5 mm overlap with the polymer encapsulant 151. The cover membranes were dried at 23 ± 2 ° C. for 30 minutes and 40 ° C. for 1.5 hours.

  0.3 μL of 5% hydrophilic polyurethane (Hydromed D640 / Hydromed D4 mixture with 80% moisture content) solution in 96% EtOH (see Example 1) was dispensed onto the dried outer membrane.

All three types of layers 170, 180, 190 were dispensed on x, y, z planes fitted with an automatic dispenser (IVEK pump).
Example 1-Representative creatine and creatinine sensor construction Each creatine sensor and creatinine sensor includes a known amperometric sensor. FIG. 1 shows such a sensor 1 (above), which is a device for measuring the concentration of an analyte in a biological sample, such as ABL ^™ 735 Blood Gas Analyzer (Radiometer Medical ApS, Copenhagen, Denmark). Suitable for mounting on.

  FIG. 2 shows four layers: a noise reducing hydrous spacer layer 22 facing the platinum anode 4 of the electrode 2; an interference limiting membrane layer 23; a gasket 24 surrounding the enzyme layer 25; and a moisture content of about 80%. 3 shows details of a membrane 21 comprising a diffusion-limited porous polymeric material 26 impregnated with a hydrophilic polyurethane having The coated membrane layer 26 faces the sample to be analyzed.

The spacer layer 22 is a 21 ± 2 μm track-etched film of polyethylene terephthalate (PETP) (pore diameter: about 1.3 to 1.5 μm; pore density: 2.2 · 10 ⁷ pores / cm ² ). It may be. The interference limiting membrane layer 23 may be a 6 ± 2 μm porous membrane of cellulose acetate (CA).

  The gasket 24 may be a 30 ± 5 μm double-sided adhesive disk with a 1500 μm diameter center hole. The adhesive of the gasket 24 adheres to the interference restriction layer 23 and the diffusion restriction layer 26 to such an extent that the enzyme is not leaked from these layers.

  The enzyme layer 25 of the creatine sensor is typically an approximately 20 μm layer of creatinase and sarcosine oxidase cross-linked to glutaraldehyde mixed with a suitable additive such as a buffer. The enzyme layer 25 of the creatinine sensor is typically an approximately 20 μm layer of creatininase, creatinase and sarcosine oxidase crosslinked to glutaraldehyde mixed with a suitable additive such as a buffer.

The diffusion-limited porous polymeric material 26 is an about 12 μm layer (pore diameter about 0.1 μm) of polyethylene terephthalate (PETP) impregnated with hydrophilic polyurethane (Hydromed D640 / Hydromed D4 mixture having a water content of about 80%); Pore density: 3 · 10 ⁷ pores / cm ² ) (see Example 1).

In the case of a creatinine sensor, both creatine and creatinine are converted to hydrogen peroxide. In the case of a creatine sensor, only creatine is converted into hydrogen peroxide.
At the amperometric electrode, hydrogen peroxide is anodized at +675 mV against Ag / AgCl. The resulting current flow is proportional to the creatinine / creatine concentration in the sample.

  The concentration of creatinine is determined from the difference between the creatinine sensor signal (indicating creatine + creatinine) and the creatine sensor signal (indicating creatine).

Example 2-Relative diffusion coefficient of imidazole / H-imidazole ⁺ through a cellulose acetate membrane The relative diffusion coefficient of imidazole / H-imidazole ⁺ through a CA membrane is in contact with a solution buffered by imidazole through a CA membrane. It was measured by analyzing the pH in the unbuffered solution obtained. Both samples were 30 mL and the CA membrane in contact with both solutions was 10 mm in diameter (see “Measurement of Apparent Diffusion Coefficient” above). The solution was 140 mM NaCl, 4 mM KCl, 1 mM CaCl ₂ and 110 mM imidazole, and the pH was adjusted to 7.40 at 25 ° C. After 20 hours, the pH of the unbuffered solution rose to 8.86 at 25 ° C., while the pH of the other solution was unchanged. Thus, the neutral species of imidazole (the basic part of the imidazole pair) is able to diffuse at least 30 times faster than the charged species, reflecting a significant pH change, but it can be applied to the calibration solution and the sample respectively. It can be observed on the electrode surface during the exposure period.

Example 3-Influence of a hydrous spacer layer on creatinine sensor measurements
Two similar hematology analyzers of type ABL ^™ 735 from Radiometer Medical ApS, Denmark were modified to accommodate the dual sensor system described in Example 1. In one blood analyzer, five creatinine sensors (three enzyme sensors) and five creatine sensors (two enzyme sensors) each equipped with a spacer layer according to Example 1 were arranged. Similarly, in another hematology analyzer, five creatinine sensors (three enzyme sensors) and five creatine sensors (two enzyme sensors) were arranged according to Example 1, but without a spacer layer.

  A calibration solution was prepared by dissolving 200 μM creatinine in Radiometer Calibration Solution 1 S1720 and 200 μM creatine in Radiometer Calibration Solution 2 S1730. These solutions were repeatedly introduced into one apparatus after other and sensor responses were obtained.

  After such calibration, the sensors were subjected to the following seven creatinine-free and creatine-free solutions to measure their “pseudo” creatine responses (see Table A).

  The sensor signal is calculated as the difference between the signal immediately before subjecting the sensor to the sample and the signal 28 seconds after contact with the sample. Similarly, sensor sensitivity is calculated from a calibration solution. The sensor signals for the seven liquids are then converted to the appropriate “pseudo” creatine concentrations, and both 2-enzyme and 3-enzyme sensors are compared using sensor sensitivity (Tables B and C and (See FIG. 4). It is noted that the expected value is “0” (zero).

  As can be seen from Tables B and C, the sensor equipped with the spacer layer has a slightly different difference ("pseudo" creatine with 2 enzyme sensors and 3 enzyme sensors respectively) compared to the sensor without the water-containing spacer layer. Difference between signals) and a smaller standard deviation.

  A series of similar experiments were performed with a 3-enzyme sensor with and without a hydrous spacer layer. These results are in this case converted into “pseudo” creatinine signals. The results are shown in FIG. 4 (the black diamond is a sensor with a spacer layer and the blank square is a sensor without a spacer layer). Sensor-to-sensor variation is reduced with the introduction of a spacer layer.

Example 4-Effect of Water-containing Spacer Layer on Whole Blood Creatine / Creatinine Sensor Measurements Whole blood samples from healthy individuals were measured with the above sensor according to Example 1. The creatinine concentration of the patient blood sample was calculated using all combinations of 5 creatine sensors and 4 creatinine sensors from each of the top two lines (with and without spacer layer) according to Example 3 (Table See D and Table E).

As can be seen from Tables D and E, the span and standard deviation are reduced for sensors that include a spacer layer.
Sensitivity platinum electrode to Example 5-H ₂ O _2, the three simple planar sensors only include water spacer layer and interference reduction layer, in essence, "General sensor constructed above (thick film sensor Type) ”.

  A hydrous porous spacer layer is formed by dispensing 300 nL of a 7% D4 PUR (Hydromed inc.) Solution in 96% ethanol on a Pt working electrode by micro-dispensing and 1 μm thick cellulose acetate and cellulose An acetate butyrate inner membrane was fabricated on top of the spacer layer.

Two reference flat sensors were manufactured as described above, but without the spacer layer and with an inner membrane thickness of 2 μm.
A series of measurements was performed using 5 sensors and 1 mM H ₂ O ₂ solution as the test sample. The H ₂ O ₂ solution was used to simulate a glucose oxidase layer exposed to glucose. The results are shown in FIG.

Higher sensitivity is observed for the sensor with the spacer layer (about 55 nA) than for the sensor without the spacer layer (about 35 nA). This is believed to be the result of improved H ₂ O ₂ diffusion to the platinum electrode because the spacer allows for a high diffusion rate and the cellulose acetate inner membrane is the spacer This is because, if a layer is present, it can be distributed in a very thin layer.

FIG. 1 shows a conventional enzyme sensor comprising an electrode and a membrane. FIG. 2 shows the membrane of the sensor of FIG. FIG. 3 shows a typical planar thick film sensor construction. FIG. 4 shows the effect of introducing a hydrous spacer layer on the pseudo signal for a dual sensor system in contact with various liquids. FIG. 5 shows the sensitivity to H ₂ O _{2 of} a sensor coated with a hydrous spacer layer and an interference exclusion layer.

Claims (20)

An amperometric enzyme sensor for measuring the concentration of an analyte in a fluid sample,
electrode,
A hydrous spacer layer in contact with the electrode;
At least one intermediate layer, the innermost part of the intermediate layer being in contact with the spacer layer, and at least one enzyme layer, wherein the innermost part of the enzyme layer is the innermost part of the intermediate layer Amperometric enzyme sensors, including those in contact with the outside.   The spacer layer has a porosity in the range of 0.0005-2% (vol / vol) for the track-etched material and a porosity in the range of 1-90% for the solution cast material. The enzyme sensor described in 1.   The enzyme sensor according to claim 1 or 2, wherein the spacer layer comprises water and solids, and the weight ratio between water and solids is in the range of 10: 1 to 1:10.   The enzyme sensor according to any one of claims 1 to 3, wherein the spacer layer has a thickness in the range of 0.2 to 20 µm.   The enzyme sensor according to any one of claims 1 to 4, wherein the water-containing spacer layer further comprises one or more components selected from a buffer, an electrolyte salt, and a cation exchange material.   The sensor is a conventional sensor, and the spacer layer comprises a solid material consisting essentially of a porous polymer matrix selected from polyethylene terephthalate (PETP), polyvinyl chloride and polycarbonate; The enzyme sensor according to any one of claims 1 to 5.   The enzyme sensor according to claim 6, wherein the porous polymer matrix is a polyethylene terephthalate (PETP) material.   The enzyme sensor according to any one of claims 1 to 7, wherein the porous polymer base material of the spacer layer is a track-etched film. The sensor is a flat sensor type and the spacer layer is solid and is hydrophilic polyurethane, hydrophilic poly (meth) acrylate, poly (vinyl pyrrolidone), polyurethane, Nafion ^™ polymer, electropolymerized polymer and SPEES- The enzyme sensor according to any one of claims 1 to 5, comprising a solid consisting essentially of a porous polymer matrix selected from PES.   The enzyme sensor according to any one of claims 1 to 7 and 9, wherein the porous polymer base material of the spacer layer is a solution casting layer. The enzyme sensor according to any one of claims 1 to 10,
electrode,
A hydrous spacer layer in contact with the electrode;
An enzyme sensor comprising: an intermediate layer in contact with the spacer layer; and an enzyme layer in contact with the intermediate layer.   The enzyme sensor according to any one of claims 1 to 11, wherein the enzyme layer contains creatinase and / or creatininase.   The enzyme sensor according to claim 1, further comprising a cover film for the enzyme layer.   The cover membrane includes at least one porous polymeric material, wherein the outer surface and pore opening of at least one side of the at least one porous polymeric material are hydrophilic polyurethane and hydrophilic poly (meta 14. The enzyme sensor according to claim 13, which is coated with a hydrophilic polymer selected from acrylates. An amperometric enzyme sensor for measuring the concentration of creatine in a fluid sample,
Metal electrodes,
A hydrous spacer layer in contact with the metal electrode;
An interference limiting layer in contact with the spacer layer;
An enzyme layer comprising sarcosine oxidase and creatinase in contact with the interference limiting layer, and a cover membrane layer for the enzyme layer, wherein the cover membrane layer comprises a porous polyethylene terephthalate material; and Wherein the outer surface of the porous polyethylene terephthalate material and the pore opening of the porous polyethylene terephthalate material comprise a polyethylene glycol main chain segment in a weight ratio of at least 5% (w / w) polyethylene glycol segment. An amperometric enzyme sensor coated with and / or having a wet moisture content of at least 25% (w / w).   The enzyme sensor according to claim 15, wherein the porous polyethylene terephthalate material is a track-etched material. An amperometric enzyme sensor that measures the concentration of creatinine in a fluid sample,
Metal electrodes,
A hydrous spacer layer in contact with the metal electrode;
An interference limiting layer in contact with the spacer layer;
An enzyme layer comprising sarcosine oxidase, creatininase and creatinase in contact with the interference limiting layer, and a cover membrane layer for the enzyme layer, wherein the cover membrane layer comprises a porous polyethylene terephthalate material And wherein the outer surface of the at least one face of the porous polyethylene terephthalate material and the pore opening comprise at least 5% (w / w) weight ratio of polyethylene glycol segments to polyethylene glycol main chain segments. An amperometric enzyme sensor coated with a hydrophilic polyurethane and / or having a wet moisture content of at least 25% (w / w).   The enzyme sensor according to claim 17, wherein the porous polyethylene terephthalate material is a track-etched material.   A device for measuring the concentration of an analyte in a fluid sample, comprising one or more enzyme sensors according to any one of claims 1-18. A method for measuring the concentration of an analyte in a fluid sample,
A method comprising contacting the fluid sample with the enzyme sensor according to any one of claims 1 to 18, and performing at least one measurement requiring an electrode of the enzyme sensor.

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