Electrode system for measuring analyte concentration under in vivo conditionsTechnical Field
The invention relates to an electrode system for measuring an analyte concentration under in vivo conditions, having the features specified in the preamble of claim 1. An electrode system of this type is known from WO 2007/147475.
Background
Sensors with implantable or insertable electrode systems facilitate measurements of physiologically significant analytes such as, for example, lactate or glucose in body tissue of a patient. The working electrode of this type of system has an electrically conductive enzyme layer in which enzyme molecules are bound, which release charge carriers by catalytic conversion of analyte molecules. In this process, a current is generated as a measurement signal, the amplitude of which is related to the analyte concentration.
The conductivity of the good enzyme layer should be as high as possible so that the released charge carriers are detected as completely as possible as measurement signal; these enzyme layers should have sufficient water permeability to allow diffusion of analyte molecules from aqueous body fluids, typically interstitial fluid or blood, to the enzyme layer; and finally, the enzyme layers should bind the enzyme molecules contained therein as completely as possible so that they do not leak into the surrounding body tissue.
Suitable enzyme layers can be made, for example, from platinum black (which, owing to its sponge-like structure, can be impregnated with the enzyme solution and exhibits good water permeability) or from electrically conductive particles (for example carbon or metal particles) and a binder. Enzyme layers of this type are often fragile. Thus, the enzyme layer of the working electrode of the known sensor covers only a very small area, typically only a fraction of a square millimeter. An example of this is the electrode system known from US 4,655,880, in which the conductor of the working electrode extends only about 200 μm. In order to increase the local current density, the conductor is provided with an electrically insulating coating in which small openings of about 10 μm in diameter are etched before the entire length of the conductor is coated with an enzyme-containing paste in order to form an enzyme layer.
However, despite extensive research and development, known electrode systems are susceptible to interference and attendant disadvantages, as they can only be used to determine analyte concentrations with less accuracy and reliability than can be achieved using conventional in vitro assays.
To improve the accuracy of the measurement, US 2005/0059871a1 proposes the simultaneous use of multiple working electrodes to measure the concentration of the analyte of interest, and the statistical analysis of the measurement results thus obtained. As for the additional measurements, it is proposed to use other sensors to determine other analyte concentrations or physiological parameters, and to test the plausibility of the individual results on the basis of the concentrations of the different analytes thus obtained.
However, the use of a large number of working electrodes not only increases the use of device resources, but also causes problems if the inconsistent measurement signals of the individual working electrodes make it unclear which of the different measurement values accurately reflects the analyte concentration in the patient.
Disclosure of Invention
It is therefore an object of the present invention to create a method for more reliable and accurate measurement of analyte concentrations in the human or animal body.
This object is achieved by an electrode system for in vivo measurement of a concentration with respect to an analyte having the features specified in claim 1. Advantageous developments of the invention are the subject matter of the dependent claims.
The invention provides an enzyme layer of a working electrode in the form of a plurality of domains (multiple fields) arranged at a distance from each other on a conductor of the working electrode. Preferably, at least two of the domains are spaced at least 3mm, preferably at least 5mm, from each other. For example, a series of multiple domains (multiple fields) may be provided, whereby the distance between the first and last domain of the series is greater than 5 mm. The individual domains of the working electrode according to the present invention essentially form a series of working electrodes arranged in sequence. Between these domains, the conductors of the working electrode may be covered with an insulating layer. By arranging the domains of the enzyme layer on top of the openings of the electrically insulating layer, the signal to noise ratio can be improved.
The invention allows a significantly more reliable measurement of the analyte concentration in the body tissue of a patient. By providing the enzyme layer of the working electrode in the form of individual domains (e.g. at least 3mm, preferably at least 5mm apart), analyte concentrations can be measured in relatively large volumes. With the electrode system according to the invention, it is thus of no importance that only the disturbing influence of cells having a small volume with a diameter of about 0.1mm is influenced. Surprisingly, most of the problems observed during in vivo measurements using known sensors seem to be based solely on the fact that: due to the small size of the enzyme layer, the analyte concentration is measured in a volume unit that is too small to be often representative of the rest of the patient's body due to transient local effects.
Due to the long distance expansion of the enzyme layer of several millimeters or more, the working electrode must have flexibility so that it can adjust its shape according to the motion of the body. However, the relatively fragile nature of known conductive enzyme layer materials does not appear to make it possible to manufacture flexible working electrodes. However, according to the measurement of the present invention, that is, the enzyme layer in the form of a plurality of domains (multiple fields) arranged at a distance from each other on the conductor of the working electrode is designed so that the working electrode can be bent without the enzyme layer peeling off, despite the brittle nature of the enzyme layer material.
The reason why measurements with conventional working electrodes having an effective range (i.e. the enzyme layer) extending less than 0.2mm square are incorrect is presumably that the patient's actions may momentarily impede fluid exchange in small volume units, e.g. cells are pressed against the working electrode and interstitial fluid is replaced, or capillaries are compressed and thus blocked. Presumably, this can result in the analyte concentration in the corresponding volume element immediately surrounding the working electrode not being representative of the rest of the patient's body. These volume elements, in which the fluid exchange is momentarily impeded, appear very small, but, and usually, have a diameter of less than 1 mm. It is likely that the elastic and soft consistency of the body tissue allows the force to relax over a very short distance and thus the exchange of body fluid in a volume element of such a small size is negatively affected. Letting at least some of the domains of the enzyme layer formed in the electrode system according to the invention be distributed at a considerable distance, for example at least 5mm apart from each other, preferably at least 1cm apart from each other, it is thus provided that, even in the worst case, only small and usually negligible parts of the working electrode are negatively affected.
In this case, the distance between two domains of the enzyme layer of the electrode system according to the invention is measured from the edge of one domain to the edge of the other domain facing it.
In this case, the distance between adjacent domains is preferably at least 0.3mm, in particular at least 0.5 mm. Each of the individual domains preferably extends less than 2mm, preferably less than 1mm, in particular less than 0.6mm in two directions perpendicular to each other. These domains may be, for example, circular with a diameter of less than 1mm, or rectangular with a side of less than 1 mm. Preferably, the domains are arranged in rows on the conductor of the working electrode. However, it is also possible to arrange the domains on a circular or rectangular conductor, for example in several rows and columns. The number of these fields is virtually freely selectable. However, the working electrode preferably has at least 5 domains.
The working electrode of the electrode system according to the invention is provided with a diffusion barrier which slows down the diffusion of analytes from the body fluid surrounding the electrode system to the enzyme molecules immobilized in the enzyme layer. The diffusion barrier may be, for example, a cover layer covering the enzyme layer. However, it is also possible that diffusion-inhibiting particles are bound to the enzyme layer to act as a diffusion barrier. For example, the pores of the enzyme layer may be filled with a polymer through which analyte molecules can only slowly diffuse. Such polymers should be hydrophilic and have rapid water absorption. A diffusion barrier may advantageously be used to reduce analyte molecule consumption at the working electrode. If the patient's motion momentarily interferes with the exchange of body fluids around the enzyme layer domain of the working electrode, having a lower conversion rate of analyte molecules helps to reduce the effect of such interference. The lower the consumption of analyte, the longer it takes for the depletion effect to occur (i.e. the decrease in analyte concentration in the corresponding region due to the measurement being taken).
An advantageous development of the invention proposes that the working electrode should have a spacer arranged on the enzyme layer as seen from the conductor and provide a minimum distance between the enzyme layer and the body tissue surrounding the cells. The spacer forms a reservoir for analyte molecules. By this measure, the influence of transient disturbances of the fluid exchange around the working electrode can be further reduced. The spacer may be, for example, a layer made of a biocompatible polymer that facilitates analyte permeation. The spacer can be used to create an analyte buffer volume that supplies the enzyme domain of the working electrode. In this way it can advantageously be achieved that no significant negative influence on the measurement signal occurs even in the presence of significant disturbances in the fluid exchange in the surroundings of the enzyme layer domain during the hour and half. The spacer may also be provided as, for example, a perforated membrane (e.g., dialysis membrane) or a mesh. The spacers are preferably 3 to 30 microns thick. The spacer may be disposed on the diffusion suppressing cover layer. However, it is also possible to arrange the spacer directly on the enzyme layer. In this case, the spacer may also act as a diffusion barrier and slow down the diffusion of analyte molecules to the enzyme layer.
The spacer preferably covers the working electrode and the counter electrode and the reference electrode, if present, in the form of a continuous layer. The spacer preferably covers the entire implanted surface of the substrate. If the spacer is made of a biocompatible material, the tissue response to implantation may be reduced. Independently of this, it is also preferable that the spacer is disposed on the diffusion-suppressing cover layer and is more hydrophilic than the cover layer.
Like the electrical conductor of the counter electrode of the electrode system according to the invention, the electrical conductor of the working electrode is preferably designed as a conductor path on a substrate, for example made of metal or graphite on a plastic plate. However, the conductor may also be designed in the form of a wire. The enzyme layer domains may be disposed on a conductor designed in the form of a wire (e.g., in the form of a loop segment). In the case of a loop segment on a wire, it should be understood that the object explained hereinbefore (i.e. each individual domain extending less than 2mm, preferably less than 1mm, in particular less than 0.6mm in two directions perpendicular to each other) means that the width of the loop and its diameter are in two directions perpendicular to each other.
The use of metal leads and conductor paths on the substrate allows the design of flexible sensors that can be bent through 90 degrees or more in a patient without breaking.
It is common for the conductor paths to extend only on a single side of the substrate. In principle, however, it is also possible for a single conductor to run on both sides of the substrate, for example by drilling or on the surrounding side.
Another advantageous development of the invention is to provide an enzyme to interact with a catalytic redox mediator in the enzyme layer, wherein the catalytic redox mediator reduces or prevents the oxygen dependence of the catalytic conversion of the analyte. This type of catalytic redox mediator is sometimes also referred to as an electrocatalyst because it prefers to transfer electrons to the conductive components of the working electrode, such as graphite particles in the enzyme layer. Manganese dioxide, for example in the form of pyrolusite, or other metal oxides that catalytically oxidize hydrogen peroxide and transfer electrons to the conductive component of the working electrode in the process can also serve as catalytic redox mediators. The catalytic redox mediator in the form of a metal oxide can advantageously lower the potential of the working electrode by more than 100 mv, thereby significantly reducing the effect of interfering substances (e.g. ascorbate or uric acid) on the measurement signal. In the case of enzymes that oxidize analyte molecules and produce hydrogen peroxide in the process, the use of a catalytic redox mediator of this type makes it possible to counteract the oxygen depletion around the working electrode and thus make the rate of conversion depend only on the analyte concentration and not on the oxygen concentration over a wide range of concentrations.
Organometallic compounds (e.g., cobalt-phthalocyanine) are also suitable as catalytic redox mediators for the degradation of hydrogen peroxide. The catalytic redox mediator may be covalently bound to the enzyme molecule or embedded within the enzyme layer, e.g., in the form of separate particles.
However, the catalytic redox mediator may also produce direct electron transfer. By this method, a catalytic redox mediator covalently bound to the enzyme can be used to perform the oxidation of the analyte molecule and the transfer of electrons to the working electrode without the need for an intermediate step of hydrogen peroxide generation. In direct electron transfer, electrons from the prosthetic group of the enzyme are transferred directly to the catalytic redox mediator and from there to the conductive component of the working electrode, such as graphite or metal particles in the enzyme layer.
Direct electron transfer can be produced, for example, with enzymes such as dehydratases having pyrroloquinoline quinone (PQQ) as a prosthetic group. In the case of glucose dehydratase (GlucDH) from Acinetobacter calcoaceticus, PQQ is converted to a reduced state (reduced state) by this enzyme during glucose oxidation. The prosthetic group can be directly covalently bound to a conductor using gold nanoparticles to facilitate direct electron transfer from the reduced PQQ to the conductor. It is clear that oxygen does not react with reduced PQQ, which means that it does not compete with the electron transfer. Another way to transfer these electrons from the reduced PQQ is based on the recognition that PQQ itself can exist in multiple oxidation stages
And thus can serve as a catalytic redox mediator. Thus, additional PQQ molecules may be covalently bound to the GlucDH enzyme and used to receive electrons from the catalytically active pocket (i.e., the prosthetic group of PQQ present therein) of the protein. Another option is based on a variant of GlucDH (from Pseudomonas cepacia) having flavin-adenine-dinucleotide (FAD) as a prosthetic group and cytochrome C protein in the other subunits, which can be derived from FADH2Transfer electrons.
The electrode system according to the invention can additionally have a reference electrode. The reference electrode can supply the working electrode with a reference potential defined by, for example, a silver/silver chloride redox system. Furthermore, the electrode system according to the invention may additionally have electrodes, for example additional working electrodes, in particular working electrodes with different measurement sensitivities as described in US 2007/0151868a 1.
In combination with a potential self-regulator connected to the electrode system and an amplifier for amplifying the measurement signal, the electrode system according to the invention forms a sensor. Preferably, the amplifier and the potential adjustor are arranged on a printed circuit board carrying conductors of the counter electrode and the working electrode. The electrodes may be arranged on the substrate, for example in the form of a plastic plate with one end attached to the circuit board. The printed circuit board may also be integrated into the substrate provided with these electrodes. The potential self-regulator and the preamplifier may for example be arranged on a flexible plastic plate of the substrate which simultaneously forms the conductor paths of the printed circuit board and the electrode system.
Another aspect of the invention is directed to an electrode system for measuring an analyte concentration under in vivo conditions comprising an opposing electrode having an electrical conductor, a working electrode having an electrical conductor, the electrical conductor of the working electrode having disposed thereon an enzyme layer comprising immobilized enzyme molecules for catalytically converting the analyte, and a diffusion barrier for slowing the diffusion of the analyte from a bodily fluid surrounding the electrode system to the enzyme molecules, wherein the diffusion barrier is a layer covering the enzyme layer, characterized in that the diffusion barrier is made of a mixture of at least two polymers. The enzyme layer of such an electrode system is preferably, but not necessarily, designed in the form of multiple domains (multiple fields).
The diffusion barrier of such an electrode system is a solid solution of at least two polymers, preferably acrylates. Thus, the diffusion barrier may combine advantageous properties of different polymers in terms of permeability, water absorption, swelling and flexibility. One or all of these polymers may be acrylates. Preferably, one or all of these polymers are copolymers, especially copolymers of hydroxyethyl methacrylate. Copolymers are polymers made by polymerizing at least two different monomers. Hydroxyethyl methacrylate has been found to have excellent water absorption combined with slight swelling.
It is advantageous to use mixtures of polymers having different glass transition temperatures. For example, one polymer may have a glass transition temperature below 90 ℃, particularly below 70 ℃, whereas another polymer has a glass transition temperature above 100 ℃, particularly above 110 ℃. The glass transition temperature was measured by differential scanning calorimetry using a heating rate of 10K per minute.
For example, the diffusion barrier may be a mixture of a copolymer of methyl methacrylate and hydroxyethyl methacrylate and a copolymer of butyl methacrylate and hydroxyethyl methacrylate. Such a mixture may for example comprise 5 to 50% by weight of butyl methacrylate and hydroxyethyl methacrylate to obtain good flexibility of the diffusion barrier.
Drawings
Further details and advantages of the invention will now be described according to exemplary embodiments with reference to the accompanying drawings. In the figure:
FIG. 1: an exemplary embodiment of an electrode system according to the present invention is shown;
FIG. 2: showing a detail of figure 1;
FIG. 3: another detail view of fig. 1 is shown;
FIG. 4: showing a section along section line CC of fig. 2;
FIG. 5: displaying in vitro function curves for the electrode system shown in figure 1 with different cover layers;
FIG. 6: showing an example of in vivo measurement of the electrode system according to the present invention; and
FIG. 7: comparative in vitro measurements of the electrode system at various oxygen concentrations are shown.
Detailed Description
Fig. 1 shows an exemplary embodiment of an electrode system for insertion into human or animal body tissue, for example into skin or subcutaneous adipose tissue. An enlarged view of detail a is shown in fig. 2 and an enlarged view of detail B is shown in fig. 3. Fig. 4 shows a corresponding cross-sectional view along the section line CC of fig. 2.
The electrode system shown has a working electrode 1, a counter electrode 2, and a reference electrode 3. The conductors 1a, 2a, 3a of the electrodes are arranged on the substrate 4 in the form of metallic conductor paths, preferably made of palladium or gold. In the exemplary embodiment shown, the substrate 4 is a flexible plastic plate, for example made of polyester. The base plate 4 is less than 0.5mm thick, for example 100 to 300 microns, and is therefore easily bendable so that it can accommodate the action of the surrounding body tissue after its insertion. The base plate 4 has a narrow shaft for insertion into the tissue of the patient's body and a wide head for connection to an electrode system arranged outside the body. The axial length of the substrate 4 is preferably at least 1cm, in particular 2cm to 5 cm.
In the shown example embodiment, a part of the measuring device, i.e. the head of the base plate, protrudes from the body of the patient during use. Alternatively, it is also possible to implant the entire measuring facility and to transmit the measurement data wirelessly to a receiver arranged outside the body.
The working electrode 1 carries an enzyme layer 5 containing immobilized enzyme molecules for catalytically converting an analyte. The enzyme layer 5 may be applied in the form of a hardened paste of, for example, carbon particles, polymeric binders and enzyme molecules. Details of the manufacture of this type of enzyme layer 5 are disclosed, for example, in WO 2007/147475, which is incorporated herein by reference. The analyte to be measured may be, for example, glucose, lactate or other medically significant molecules. Typically, an oxidase is used as an enzyme, for example glucose oxidase or lactate oxidase, or a dehydratase, such as glucose dehydratase.
In the exemplary embodiment shown, the enzyme layer 5 is not applied continuously on the conductor 1a of the working electrode 1, but in the form of individual domains arranged at a distance from one another. Although the enzyme layer 5 is fragile, this allows bending of the electrode system without peeling off the enzyme layer 5. The electrode system shown can thus be bent over 90 ° without breaking, so that it can adapt to the body movements after its insertion.
The individual domains of the enzyme layer 5 in the exemplary embodiment shown are arranged in a series, whereby the distance between the first and last domain of the series is greater than 1 cm. The distance between each of the adjacent domains is at least 0.3mm, in particular more than 0.5mm, whereby the distance is measured from the edge of one domain to the edge of the other domain. The individual domains each extend 0.2mm to 0.6mm, for example 0.2mm to 0.4mm, in two directions perpendicular to each other. The shape of the field may be, for example, circular or square. The total area of all domains can be chosen substantially freely. Generally, a total area of less than 1 square millimeter is sufficient. The total area in the exemplary embodiment shown is about 0.4 to 0.6 square millimeters.
The conductor 1a of the working electrode 1 has a narrow position between these enzyme layer domains, which is particularly clear from fig. 2. The conductor 2a of the counter electrode 2 has a contour following the course of the conductor 1a of the working electrode 1. This results in a sandwiched or interlocked arrangement of working electrode 1 and counter electrode 2 with an advantageously short current path and low current density. The conductor 1a of the working electrode 1 of the exemplary embodiment shown is designed to be narrow and to have a width of less than 1 mm. In the exemplary embodiment shown, the width of the conductor 1a at its wide position covered by the domains of the enzyme layer 5 is less than 0.6mm, i.e. about 0.3mm to 0.5 mm. The width of the conductors 1a and 2a at the intermediate narrow position is less than 0.3mm, i.e. 0.05mm to 0.2 mm. However, these conductors are not necessarily in a sandwiched configuration. In principle, the conductors 1a, 2a can also be designed in a straight line and have a fixed width.
In order to increase its effective area, the counter electrode 2 may be provided with a hole-like conductive layer 6 in the form of individual domains on the conductor 2a of the counter electrode 2. Such as the enzyme layer 5 of the working electrode 1, this layer 6 may also be applied in the form of a hardened paste of carbon particles and a polymeric binder. The domains of this layer 6 preferably have the same size as the domains of the enzyme layer 5, but this is not essential. However, the measures for increasing the surface of the counter electrode can also be carried out beforehand, and the counter electrode 2 can also be designed as a straight conductor path without any kind of coating.
The reference electrode 3 is disposed between the conductor 1a of the working electrode 1 and the conductor 2a of the counter electrode 2. The reference electrode shown in fig. 3 is composed of a conductor 3a, and a domain 3b of conductive silver/silver chloride paste is disposed on the conductor 3 a.
Fig. 4 shows a schematic cross-sectional view along the section line CC of fig. 2. The cross-section line CC extends through one of the enzyme layer domains 5 of the working electrode 1 and is between the domains of the conductive layer 6 of the counter electrode 2. Between the domains of the enzyme layer 5, the conductor 1a of the working electrode 1 may be covered by an electrically insulating layer 7, such as the conductor 2a of the counter electrode 2, between the domains of the conductive layer 6, to prevent interfering reactions that may be catalyzed by the metal of the conductor path 1a, 2 a. Thus, the domains of the enzyme layer 5 are located within the openings of the insulating layer 7. Likewise, the domains of the conductive layer 6 of the counter electrode 2 may also be located on top of the openings of the insulating layer 7.
The enzyme layer 5 may be covered by a cover layer 8, which cover layer 8 exhibits a diffusion resistance to the analyte to be measured and thus acts as a diffusion barrier. The cover layer 8 may, for example, consist of polyurethane, acrylate, in particular a copolymer of methyl methacrylate and hydroxyethyl methacrylate, or another polymer which exhibits slight swelling but absorbs water rapidly. The cover layer 8 can advantageously be made of a mixture of at least two different acrylates, each of which can be a copolymer. Particularly preferred results are obtained by mixing a copolymer of methyl methacrylate and hydroxyethyl methacrylate with a copolymer of butyl acrylate and hydroxyethyl methacrylate.
The preferred thickness of the cover layer 8 is, for example, 3 to 30 micrometers. Due to the diffusion resistance of the cover layer 8, it results in less analyte molecules per unit time reaching the enzyme layer 5. Thus, the cover layer 8 reduces the rate at which analyte molecules are converted and thus counteracts the loss of this analyte concentration.
The covering layer 8 extends substantially continuously over the entire area of the conductor 1a of the working electrode 1. On the cover layer 8, a biocompatible membrane is arranged as spacer 9, which establishes a minimum distance between the enzyme layer 5 and the cells of the surrounding body tissue. This advantageously creates a reservoir of analyte molecules from which they can reach the corresponding enzyme layer domain 5 if a transient disturbance of the fluid exchange takes place around the enzyme layer 5. If the fluid exchange around the electrode system is momentarily limited or even prevented, the analyte molecules stored in the spacer 9 continue to diffuse to the enzyme layer 5 of the working electrode 1, which converts them. The spacers 9 therefore only lead to a significant loss of analyte concentration and a corresponding distortion of the measurement result after a significantly longer period of time. In the exemplary embodiment shown, the membrane forming the spacer 9 also covers the counter electrode 2 and the reference electrode 3.
The spacer membrane 9 may be, for example, a dialysis membrane. In this case, it is understood that the dialysis membrane is a membrane which is impermeable to molecules larger than the maximum size. The dialysis membrane can be prefabricated in a separate manufacturing process and can then be applied during the manufacture of the electrode system. The maximum size of the molecules that can penetrate the dialysis membrane is chosen so that analyte molecules can pass through while larger molecules are left behind.
Alternatively, as an alternative to a dialysis membrane, a coating made of a polymer with a high permeability to analytes and water (e.g. polyurethane based) can be applied on the electrode system as a spacer membrane 9.
The enzyme layer 5 may contain metal oxide particles, preferably manganese dioxide particles, as catalytic redox mediators. Manganese dioxide catalyzes the conversion of hydrogen peroxide, for example, formed by the enzymatic oxidation of glucose with other biological analytes. During the degradation of hydrogen peroxide, the manganese dioxide particles transfer electrons to the conductive component of the working electrode 1, e.g. graphite particles within the enzyme layer 5. The catalytic degradation of hydrogen peroxide counteracts any reduction in oxygen concentration in the enzyme layer 5. This advantageously allows the conversion of the analyte to be detected in the enzyme layer 5 to be unrestricted by the local oxygen concentration. The use of a catalytic redox mediator therefore counteracts the distortion of the measurement results due to the low oxygen concentration. Another advantage of the catalytic redox mediator is that it prevents the production of cell-damaging concentrations of hydrogen peroxide.
Fig. 5 shows the curves of the functions of the electrode system described above with different coatings 8 measured in vitro. The measured current intensity (in nA) is plotted as a function of the glucose concentration (in mg/dl). The upper function curve a is measured using an electrode system whose coated membrane 8 made of hydrophilized polyurethane has a thickness of 5 μm. For comparison, the dependence of the current on the glucose concentration of the electrode system with the coated membrane 8 is shown as a curve B below, in which the coated membrane 8 exhibits approximately twice the diffusion resistance for analyte molecules, for example due to a relatively large thickness or a small hydrophilization. The electrode system of the function curve shown in fig. 5 operates with a polarization voltage of 350 mV.
Hydrophilized Polyurethanes (HPUs) for use as coatings can be produced by polycondensation of 4, 4' -methylene-bis (cyclohexyl isocyanate) with a diol mixture. The two components of the diol mixture used to adjust the degree of hydrophilization of the polymer are polyethylene glycol (PEG, MW (molecular weight) 1000g/mol) and polypropylene glycol (PPG, MW (molecular weight) 1500 g/mol). For function curve a, the HPU coating 8 is made with a ratio of 1 to 3 PEG to PPG. For function curve B, the HPU coating 8 was fabricated at a PEG: PPG ratio of 1: 7. The cover layer 8 is about 5 μm thick in both examples.
Since the analyte concentration around the electrode system should be influenced by the measurement as little as possible and therefore should not be distorted more than to a slight extent even at a momentary disturbance of the exchange of body fluid, it is advantageous to have a low analyte conversion rate and thus a low measurement current. The total area of utilization at a glucose concentration of 180mg/dl is 1mm2Or smaller enzyme layers, produce currents of less than 50nA, especially less than 10nA, with good results obtained with electrode systems. An electrode system such as the function curve B shown in fig. 5 was used to measure a current of 3nA in pig skin at a glucose concentration of 180 mg/dl. Such small measurement signals are difficult to transmit over long distances. Therefore, it is preferred to arrange the potential self-regulator and the amplifier in close proximity to the electrode system. The potential adjuster 10 and the amplifier 11 may be disposed, for example, on the head of the substrate 4, as shown in fig. 1. It is also possible to attach the substrate 4 to a conductor path board carrying the potential self-regulator and amplifier.
Fig. 6 shows in vivo measurements in subcutaneous adipose tissue in the abdomen of an insulin dependent diabetic using two electrode systems whose function curves are shown in fig. 5 and, furthermore, they are equipped with spacers. The two electrode system is implanted at a distance of about 10 cm.
The signal characteristics of the two simultaneously implanted sensors show that the results obtained with the electrode system according to the invention are very consistent. There was no relevant difference in local glucose concentration between the two insertion sites. The results shown also demonstrate that there appears to be no instantaneous difference in glucose concentration between blood (circles) and tissue (solid and dashed lines).
The current values of the two sensors are converted to glucose values by calculation using one measurement per minute and a sample rate without filtering.
The conversion is based on the blood glucose level measured on a body fluid sample under in vitro conditions.
For the in vivo measurement shown in fig. 6, the electrode system was first coated with a hydrophilic polyurethane coating 8 and then immersed in a 12.5% ethanol solution of Butyl Methacrylate (BMA) and 2-Methacryloyloxyethyl Phosphorylcholine (MPC) (Lipidure CM5206, NOF Corp, japan), and then the coating thus produced was dried for 12 hours to a thickness of 25 μm. The current density was not substantially changed by the spacer 9 made by BMA-MPC: the function curve A of the electrode system of FIG. 5 without spacer 9 reaches approximately 40nA/mm at 180mg/dl2However, it reached 38nA/mm in the presence of BMA-MPC2. No difference in current amplitude was detected in the electrode system example of function curve B: 10nA/mm at 180mg/dl in the presence and absence of spacer 9 made from BMA-MPC2。
The spacer suppresses the influence of in vivo action on the sensor. Thus, the portion of the amplitude contributing to the sensor signal fluctuations in this example embodiment that is significantly related to the effect of motion is reduced from 5 to 25% of the average signal height to 0.5 to 5% of the average signal height by the spacers.
FIG. 7 shows a bar graph of the current I measured under in vitro conditions for three different glucose concentrations g (i.e., 0mg/dl, 180mg/dl and 360mg/dl) for each of two different oxygen concentrations, 0.22mmol/l (left bar of each case) and 0.04mmol/l (right bar of the paired bars shown in each case), respectively. Measurements were performed on the electrode system described above, whereby the enzyme layer 5 was built up in order to ensure direct electron transfer. GlucDH (EC 1.1.99.17) from Acinetobacter calcoaceticus (Acinetobacter calcoaceticus) was used as the enzyme. In a first step, additional PQQ molecules are first covalently bound to GlucDH as a catalytic redox mediator, for example by adding the enzyme to PQQ acid chloride. In a second step, carbon nanotubes (NanoLab, Newton, MA, USA; multi-walled CNTs, research grade) are added to a graphite-containing paste in order to improve conductivity and porosity, the paste is then mixed with the PQQ-modified GlucDH, and the working electrode paste so produced is printed onto the conductor paths 1a in a dispersed configuration and then hardened under vacuum at 40 ℃ for 4 hours. The electrode system is provided beforehand with an insulating layer 7, a reference electrode 3 and a counter electrode 2 with a conducting layer 6. The non-immobilized enzyme is removed by washing with a phosphate buffer. The graphite-containing paste contains a polymeric binder, for example based on polyvinyl chloride.
The cover layer 8 made of hydrophilic polyurethane (HPU, ratio of polyethylene glycol to polypropylene glycol 1: 3) was dispensed 3 times on the thus-produced enzyme layer 5 in the form of a 2.5% ethanol solution and dried at room temperature for 24 hours. The thickness of the thus-produced coating layer was 2 μm. To measure this in vitro function, the electrode system was operated at various oxygen concentrations in the glucose measurement solution and with a polarization voltage of 200mV relative to the Ag/AgCl reference electrode. The mean and standard deviation of the measured current were calculated for each of the 4 sensors. FIG. 7 shows the normal oxygen saturation value of the measured solution at about 0.22mmol/l with a significantly reduced oxygen concentration value of 0.04 mmol/l. No relevant or significant effect of the oxygen concentration on the in vitro function of the electrode system with direct electron transfer was observed.
List of reference marks
1: working electrode
1 a: conductor of working electrode
2: counter electrode
2 a: electric conductor of opposite electrode
3: reference electrode
3 a: electric conductor of reference electrode
3 b: silver/silver chloride layer
4: substrate
5: enzyme layer
6: conductive layer
7: insulating layer
8: covering layer
9: spacer
10: electric potential self-regulator
11: amplifier with a high-frequency amplifier