WO2023174485A1 - Biocapteur implantable - Google Patents

Biocapteur implantable Download PDF

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Publication number
WO2023174485A1
WO2023174485A1 PCT/DE2023/100195 DE2023100195W WO2023174485A1 WO 2023174485 A1 WO2023174485 A1 WO 2023174485A1 DE 2023100195 W DE2023100195 W DE 2023100195W WO 2023174485 A1 WO2023174485 A1 WO 2023174485A1
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potential
biosensor
analyte
measurement
sensor
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PCT/DE2023/100195
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German (de)
English (en)
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Wolfgang Schuhmann
Tim BOBROWSKI
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RUHR-UNIVERSITäT BOCHUM
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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/145—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue
    • A61B5/1468—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue using chemical or electrochemical methods, e.g. by polarographic means
    • A61B5/1486—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue using chemical or electrochemical methods, e.g. by polarographic means using enzyme electrodes, e.g. with immobilised oxidase
    • A61B5/14865—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue using chemical or electrochemical methods, e.g. by polarographic means using enzyme electrodes, e.g. with immobilised oxidase invasive, e.g. introduced into the body by a catheter or needle or using implanted sensors
    • C—CHEMISTRY; METALLURGY
    • C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
    • C12Q—MEASURING OR TESTING PROCESSES INVOLVING ENZYMES, NUCLEIC ACIDS OR MICROORGANISMS; COMPOSITIONS OR TEST PAPERS THEREFOR; PROCESSES OF PREPARING SUCH COMPOSITIONS; CONDITION-RESPONSIVE CONTROL IN MICROBIOLOGICAL OR ENZYMOLOGICAL PROCESSES
    • C12Q1/00—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions
    • C12Q1/001—Enzyme electrodes
    • C12Q1/004—Enzyme electrodes mediator-assisted
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N27/00—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means
    • G01N27/26—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means by investigating electrochemical variables; by using electrolysis or electrophoresis
    • G01N27/28—Electrolytic cell components
    • G01N27/30—Electrodes, e.g. test electrodes; Half-cells
    • G01N27/327—Biochemical electrodes, e.g. electrical or mechanical details for in vitro measurements
    • G01N27/3271—Amperometric enzyme electrodes for analytes in body fluids, e.g. glucose in blood
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/145—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue
    • A61B5/14532—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue for measuring glucose, e.g. by tissue impedance measurement

Definitions

  • the invention relates to an implantable biosensor and a method for automated continuous analyte measurement with minimal disruption of the concentration balance, comprising a microelectrode and a biocatalytically active film localized thereon, which adheres to the surface of the microelectrode by means of a crosslinker, the biocatalytic film being a redox polymer and a includes an oxygen-insensitive enzyme that has a catalytic center, as well as the use of such a sensor for minimally invasive implantation in the field of endocrinology.
  • Diabetes mellitus is a disease that is widespread worldwide and affects more than 420 million people worldwide, with an increasing incidence. It is based on a metabolic disorder in the regulation of blood sugar levels, which is due to an absolute or relative lack of insulin or insulin resistance. In order to be able to detect and treat high and low blood sugar levels in a timely manner, those affected depend on their blood sugar levels being monitored as closely as possible.
  • glucose self-monitoring GSM is well accepted and represents a major improvement in the quality of life for diabetes patients, it is significantly limited by the usual number of tests that have to be carried out manually every day.
  • CGM automated continuous glucose monitoring
  • a biosensor to be implanted for a long time primarily requires a high level of stability of all its components as well as a high level of accuracy, accompanied by a very simple and, if possible, infrequent manual calibration of the sensor. These requirements are influenced by the overall biocompatibility of the designed sensor.
  • the human body has a complex and strong self-defense mechanism when a foreign body is introduced. As a result of the implantation process, the body tissue suffers initial damage, which triggers a host reaction, which, in addition to wound healing, includes, among other things, the adsorption of plasma proteins to the foreign body. Followinged by a cellular reaction and attack by lymphocytes and macrophages, this irreversibly leads to fibrous encapsulation of the object.
  • This foreign body response persists for days to weeks and leads to a drastic change in the conditions in the immediate vicinity of the sensor. This leads to an increasing hindrance of substrate diffusion towards and - in the case of a catalytically converting sensor - product diffusion away from the actual sensor layer.
  • the FBR is working thus directly affects the measured sensor signal.
  • the partial prevention of product diffusion away from the sensor layer limits the overall stability of the sensor, as substrate conversion potentially produces by-products such as, in the case of glucose sensors, gluconic acid or reactive peroxide species, which can lead to local pH changes or radical reactions.
  • coatings with a wide variety of materials e.g.
  • the CGM systems currently approved for use on the market use two different general measurement methods, enzyme-based catalytic sensors and those based on affinity interactions on hydrogels. Both types have different advantages and disadvantages for the (long-term) determination of glucose, which are briefly explained below.
  • hydrogels modified with boronic acids are used as a detection matrix.
  • Glucose binds reversibly to the boronates and modulates the fluorescence of the surface matrix.
  • this modulation can be recorded via comparative measurements using band-filtered photodiodes and evaluated electronically in code.
  • the reversible binding with no change in the glucose molecule has the great advantage that the concentration inside the sensor matrix is in equilibrium with the external substrate concentration. Changes in the concentration in the area surrounding the sensor are compensated for by diffusion after a short time. Assuming that all components built into the sensor are generally sufficiently stable, the glucose-dependent sensor signal is only slightly influenced even after the sensor has been encapsulated for a long period of time. This means that such a sensor is in principle robust and can be used for reliable glucose determination for a long time. However, the reversible type of glucose measurement can also have significant disadvantages.
  • the glucose binding constant of the detection matrix must be precisely tailored to the prevailing conditions.
  • the substrate molecule only detaches very slowly from the detection matrix and changes in the glucose concentration in the sensor environment can only be detected very slowly by the sensor.
  • the binding constant is too weak, the accuracy of the detection is impaired.
  • molecules that are structurally very similar to glucose such as other sugar molecules or vicinal dialcohols, can also bind to the boronic acids because, unlike enzymes, they do not have any specific substrate recognition features. This means that the sensor can deliver distorted results, for example when certain medications are administered. Food additives such as mannitol or sorbitol, if they enter the blood, can also influence the measured sensor signal due to non-specific substrate recognition.
  • catalytic sensors chemically convert the substrate molecules that diffuse into the sensor layer into products.
  • the substrate binds precisely to the active site of an enzyme and forms an enzyme-substrate complex. From this, the reaction takes place with the formation of the product, which then leaves the binding pocket of the enzyme, since the binding strength to the substrate is significantly higher than to the product of the reaction.
  • the enzyme molecule is therefore available unchanged for a further catalytic reaction.
  • the enzyme essentially cleans the sensor matrix itself and is therefore quickly ready for further conversion reactions. Changes in glucose concentrations in the area surrounding such a sensor can be recorded much more quickly than with sensors based on affinity interactions.
  • Another advantage of enzyme-based sensors over those using boronic acids is the enzymes' unsurpassed substrate specificity. Through appropriate By modifying the enzymes, the specificity and reactivity for certain substrates can be further adjusted and undesirable side reaction pathways, such as with oxygen dissolved in the blood, can also be prevented. Enzymatic sensors are therefore faster, more precise and more sensitive than most competing technologies.
  • the advantage of catalytic conversion and the associated self-cleaning of the sensor matrix is also the most significant disadvantage of conventional electrochemical biosensors when they are to be used in implantable systems. Due to the fact that substrate is constantly being converted and thus “consumed” in the active sensor film, the system is not in equilibrium with its environment. The large differences in concentration between the sensor and the environment create a steep diffusion gradient.
  • the current response of an electrochemical sensor is limited, among other things, by the diffusion of the substrate to the electrode. If, in the case of continuous fibrous encapsulation, a growing diffusion barrier is formed around the sensor, the signal from the sensor can change significantly at the same external substrate concentration. This means that such a sensor would have to be constantly calibrated to the changing situation. Furthermore, the formation of a high diffusion barrier can result in a reduction in the long-term stability of the sensor, since the products and by-products of the enzymatic conversion accumulate in the sensor film and, for example, can cause local pH value changes.
  • EP 0778897 A1 describes an implantable glucose biosensor in which enzymes immobilized on a redox polymer electrode are present, the signals of which are periodically recalibrated. Further analogous systems with immobilized enzymes are mentioned in the publications EP 1230249 A1 and EP 0958495 A1.
  • a method for determining the thickness of a polymer layer of an implantable biosensor which includes an electrode and is surrounded by a polymer layer, the layer containing an enzyme and by bringing the biosensor into contact with a solution and applying a Potential on the biosensor, an oxidation or reduction of one redox-active species in the solution and a current is generated, the applied potential being switched off again after current flow and the subsequent drop in the potential being observed in order to obtain a large number of potential (V)-versus-time (t) data points, and determine the slope k of a graph of V versus 1/t, where the slope provides an indication of the thickness of the polymer layer.
  • V potential-versus-time
  • a sensor arrangement for detecting at least one analyte such as glucose comprising arrays with closable wells which can have a plurality of sensor elements which are adapted with membranes to be able to connect with the analyte.
  • Embodiments represent, for example, long-term analysis sensors such as those used in subcutaneous or transcutaneous monitoring of blood sugar levels in diabetics.
  • DE 69533260 T2 relates to in vivo enzyme biosensors, more precisely miniature glucose sensors with single-point calibration for subcutaneous glucose measurement
  • EP 1119637 B2 in particular points out the different diffusion coefficients in different media.
  • concentration of the analyte is determined using a coulometric method in which the charge is directly correlated with the amount of analyte in a defined volume.
  • time required for such a measurement to be carried out is relatively long and it is difficult to calculate defined volumes in implantable systems.
  • the object of the invention is therefore to develop a catalytic biosensor for analyte measurements with minimal disruption of the concentration balance of the analyte in order to achieve the main advantages (sensitive, accurate, substrate diffusion-independent and rapid analyte measurements over a long period of time) of the technologies on the market to use and at the same time reduce the disadvantages of the state of the art to a minimum.
  • the (re)adjustment of a substrate concentration balance between the sensor and the environment can only take place if the enzymatic reaction on the sensor can be stopped quickly, completely and reversibly. Only then will there be no permanent diffusion gradient between the compartments. This on and off process is achieved by choosing a suitable sensor architecture in which the enzyme ensemble is efficiently electrically connected to the electrode via a redox polymer.
  • the redox polymer takes on the function of a suitable, stable immobilization matrix for the enzymes and, on the other hand, it is responsible for the electron transfer between the enzyme and the electrode, which takes place via the mediators bound to the redox polymer.
  • the spatial proximity between the mediator complexes and the enzymes in the immobilized film makes it possible to quickly control the processes that occur within the catalytically active layer. This means that the enzymatic conversion process can be switched between 'on' and 'off' states quickly and reliably.
  • the enzymatic reaction can only continue if a sufficient overpotential is applied to the electrode, which enables the reoxidation of the electron-transferring Os complex.
  • an applied potential that is sufficiently more negative than the midpoint potential of the Os 2 Os 3+ pair, the mediator and ultimately also the enzyme itself is forced into a reduced state so that no (further) substrate turnover takes place; this prevents the establishment of a substrate gradient.
  • the applied potential according to the Nernst equation must be 236 mV more negative than the potential of the complex used under standard conditions.
  • the direct reaction of oxygen with Os complexes leads to two effects that must be suppressed: During glucose turnover, the generated current signal would be reduced because the reaction of oxygen with the metal complexes partially prevents efficient mediated electron transfer between enzyme and electrode. In addition, the presence of oxygen in the resting state leads to the oxidation of Os 2+ , which subsequently merges with the surrounding Enzymes can react and thereby enable undesirable substrate turnover. Therefore, the redox potential of the polymer used should be more positive than the aforementioned value, but at the same time low enough to avoid co-oxidation of interfering compounds, such as ascorbic acid or uric acid, which may be present in the analyte. In addition, the enzyme itself must be insensitive to oxygen.
  • electrochemical probes in the micrometer and nanometer range offer several advantages due to their small size, such as lower capacitance, fast spherical diffusion and negligible ohmic loss effects. These properties are of great importance for electrochemical measurements on very short time scales.
  • the redox hydrogel layer immobilized on the electrode surface must also be sufficiently thin, since electron transport in such films can only take place efficiently and without interference over short distances. Polymer films that are too thick and have inefficient electron transport can lead to inhomogeneity of charges, i.e. gradients between reduced and oxidized mediator complexes, within the film and would therefore impair or slow down the measurement.
  • an enzymatic glucose sensor shows the feasibility of the developed sensor concept. This is a combined measurement sequence that can be repeated in cycles and consists of a potentiostatic control of the sensor status between “On” and “Off” and a Data acquisition during a short potentiometric analyte measurement.
  • the universal measurement strategy presented here can also be used for substrates other than glucose (e.g. for determining glutamate or lactate) by replacing the redox-active enzymes immobilized in the hydrogel film.
  • two voltammetric cycles were carried out with a voltage feed rate of 25 mV/s in a potential range of 0.2 to 1.6 V vs. Ag/AgCl/3M KCl in a solution with 4.7 mM DAH in 0.1 M LiCIOVEtOHabs.
  • the electrodes were then freed from residues of the DAH solution alternately by dipping them twice in ethanol and distilled water. This process enables covalent bonds between the polymer backbone to be immobilized and the modified electrode surface and thus leads to increased stability of the hydrogel film on the electrode surface.
  • the amino-modified electrodes were then immersed in an aqueous solution of the bifunctional diepoxy crosslinker poly(ethylene glycol) diglycidyl ether (PEGDGE) (5 mg/mL) for 10 min to form an epoxy-terminated surface on which the free amino groups of the redox polymer backbone (structure see Fig. 1 , left - predominantly the N atom of the imidazole) can bind covalently.
  • PEGDGE poly(ethylene glycol) diglycidyl ether
  • BDGE butyl diglycol ether
  • the electrodes were coated alternately by immersion in redox polymer and a mixture of the same polymer and the oxygen-insensitive enzyme GcFAD-GDH, expressed in Pichia Pastoris, in a ratio of 1:1 (all solutions 2.5 mg/mL in water).
  • the electrodes were immersed in the pure polymer solution for 20 s, then in air with the modified tip for 30 s dried pointing below and then immersed in the polymer/enzyme solution for 60 s and air dried again for 30 s.
  • the immersion procedures were repeated several times, and the electrode was finally coated with a layer of pure polymer for 20 s.
  • the electrodes can optionally be coated with a further protective layer, for example consisting of the polymer P-(SS-BA-GMA) (structure Fig. 1 on the right).
  • the developed sensor concept is based on potentiostatic control of the potential in order to keep the sensor in substrate concentration equilibrium by switching off and preventing enzymatic glucose conversion.
  • a sufficiently negative potential of -250 mV is applied compared to the redox potential of the Os 2+ /Os 3+ transition of the mediator used. This means that the charge balance is strongly on the side of the reduced Os 2+ complex, which cannot accept electrons from the reduced enzyme.
  • a short potentiostatic charging pulse leads to an excess of Os 3+ complexes in the hydrogel film by applying a potential that is 250 mV higher than the Os 2+ /Os 3+ transition. This enables the transfer of electrons from the enzyme to the mediators and thus the enzymatic conversion of the glucose, which is located within the film in the immediate vicinity of the enzymes. This pulse must last long enough to sufficiently charge the film. At the same time, the charging process must be short enough so as not to use up too much glucose in the film. Otherwise, the substrate would diffuse from outside the film into it and the undesirable diffusion gradient would form.
  • the actual measurement of the analyte and data acquisition takes place directly after the potentiostatic pulse, whereby the circuit to the working electrode is opened and the flow of electrons is thus prevented.
  • a potentiometer with a high-resistance connected in parallel is used to measure a change in potential at the working electrode (Fig. 2).
  • the measurement sequence presented here can be carried out both analogue and on digital potentiostats with correspondingly fast data acquisition.
  • Many modern (digital) potentiostats combine the possibility of potentiometric and potentiostatic control by “internally” connecting these components in one device.
  • the measured potential change (if a suitable enzyme-Z polymer component is selected) bination) can be attributed exclusively to the enzymatic glucose oxidation and the associated reduction of the previously externally charged Os 3+ complexes. This change occurs the faster the higher the glucose concentration in the film at the time of the pulse. Therefore, this type of measurement can be used as a reliable detection method of glucose concentration.
  • the measurement concept developed was experimentally implemented and tested using previously manufactured microbiosensors of various sizes.
  • the sensors were each placed in a measuring cell with a connected inflow and outflow control and connected to a digital potentiostat in a three-electrode system with an Ag/AgCl/3M KCl reference and a platinum counter electrode.
  • PBS with a pH of 7.4 was used as the measuring electrolyte.
  • the glucose concentrations to be examined were adjusted manually via the inlet and outlet openings on the measuring cell by introducing and removing glucose solutions of different concentrations in PBS.
  • An isolated measurement involves potentiostatic charging of the redox polymer film for one second, followed by measurement data acquisition for 2.5 s, during which no potential was applied to the working electrode.
  • the sensor was switched to the “off” state over a period of 90 s by applying a potential of -50 mV (Fig. 4 left).
  • the curves shown show a change in potential as a function of glucose concentration, which is in line with the model predictions described in detail previously.
  • the potential drops sharply at the beginning of data acquisition and then weakens significantly to a quasi-linear course with a small slope, which is mainly determined by the diffusion of glucose to the sensor and into the hydrogel film.
  • One of the central aspects for a “substrate-consuming” implantable sensor that functions reliably over a long period of time is its behavior when the diffusion barrier is increased.
  • the sensor signal should, if possible, not be dependent on a fibrous protein capsule that continuously forms around the sensor compartment in the human body.
  • a sensor was modified with an additional thickened polymer layer made of P(SS-GMA-BA) so that the diffusion of glucose into the redox-active hydrogel film was severely restricted, similar to the FBR effect.
  • the overall goal was a catalytic measurement of substrate (glucose) under quasi-equilibrium conditions between the sensor film and the sensor to establish the surrounding solution.
  • substrate glucose
  • the formation of an extended linear range from a time of 0.5 s of data acquisition in the sensor shown above suggests that the sensor signal is mainly controlled by substrate diffusion to the sensor at this point in the measurement at the latest. This would not have created any advantage over an amperometric, diffusion-limited glucose measurement with the same sensor. This means that the time for data collection in our experiments still had to be significantly shortened.
  • our approach deals with determining the different slopes of the measurement curves within the first maximum 75 ms directly after the charging pulse.
  • the respective measurement data were processed with a digital fast Fourier transformation (FFT) low-pass filter (cutoff frequency ⁇ 50 Hz) and the potential curve was plotted versus time compared to the initial value after potentiostatic charging (Fig. 6 left).
  • FFT digital fast Fourier transformation
  • the slopes of the linear measurement data curve in the range of 20-45 ms show clearly recognizable differences between a small drop in potential in the presence of no or little substrate, to a larger time-dependent drop in potential at higher glucose concentrations.
  • the sensors were examined with regard to three parameters (current, potential, slope of the potential transient) and the sensor response of these parameters was plotted against the set glucose concentration in a calibration curve (Fig .
  • a potential implantable glucose sensor should provide rapid and reliable results during these changes, helping to detect such trends early.
  • a dynamic glucose profile was measured with a sensor over a period of approximately 4 hours and the magnitudes of the slopes for each measurement were plotted with constantly changing glucose concentration (FIG. 7). Meanwhile, the glucose concentration was adjusted between 1 and 7.5 mM by manually opening the inlet and outlet valves in the measuring cell. The line drawn in the graph represents the running average of four pulsed measurements.
  • the evaluation of the slopes of the potential transients at different glucose concentrations showed a very good correlation between these measurements and comparative controls with the same sensors. It was also shown that the manufactured sensors deliver the same measurement data quality as unmodified sensors after a short regeneration time, even with an artificially enlarged diffusion barrier (simulation of the FBR effect).
  • the experimental results obtained made it possible for the first time to combine the advantages of the competing glucose sensor technologies currently dominating the market (affinity sensors and electrochemical sensors) and to minimize the previous disadvantages (slow adjustment of equilibrium, influence by interference, substrate diffusion limitation after fibrous encapsulation). to reduce. It was also shown that the electrodes can also include carbon rods with a diameter of 1000 pm and/or a length of 1000 pm.
  • the measuring principle was expanded to include graphite rods with a diameter of up to 1000 pm and/or a length of 1000 pm, in particular 350 pm.
  • Figure 8 shows a measurement of glucose on a carbon rod electrode modified with a FAD-GDH/redox polymer and demonstrates the possibility of using larger electrodes for the measuring principle according to the invention, such as butyl diglycol ether (BDGE).
  • BDGE butyl diglycol ether
  • microelectrodes can be coated with a further protective layer to form the biocatalytic film, comprising the compound or the polymer poly(MPC-GMA).
  • a further protective layer to form the biocatalytic film, comprising the compound or the polymer poly(MPC-GMA).
  • Coating with this compound or polymer alone, or a number of other layer-forming polymers, is advantageous because it carries zwitterionic groups that prevent biofouling (protein and cell deposition).
  • Figure 1 shows chemical structures of the redox polymers (4) P(VI-AH)-Os and P(SS-GMA-BA).
  • Figure 2 shows the schematic measurement setup for an analog system consisting of a separate potentiostat and a potentiometer.
  • the circuit is switched externally via a computer-controlled digital in/out interface between potentiometer and potentiostat.
  • Figure 3 reveals the graphical representation of the measuring principle developed.
  • the sensor (1) is switched on or off by applying a potential of 250 mV positive or negative compared to the E1/2 of the Os complex used on the redox polymer (4).
  • a higher-level circuit to the microelectrode (2) of the sensor (1 ) enables switching between potentiostatic control and potentiometric measurement.
  • Figure 4 shows left: Isolated potential transients of a sensor (1) under changing glucose concentrations of 0 mM (PBS, pH 7.4), 1 mM, 3 mM, 5 mM and 7.5 mM (from top to bottom). Right: For each measured concentration of 0 mM or 3 mM, 4 individual potential transients were applied, which almost completely overlap each other.
  • Figure 5 shows comparative measurements of a sensor (1) which was equipped with a high diffusion barrier for the diffusion of the analyte (8), here glucose, into the active polymer film (3).
  • the longer the regeneration time the lower the depletion of the analyte (8) in the film ( 3).
  • Figure 6 shows left: Plot of the slopes of individual potential transients compared to their initial potential with increasing glucose concentration (0 mM, 1 mM, 3 mM, 5 mM and 7.5 mM - from top to bottom) with determination of the respective straight line slope in the range between 20 and 45 ms.
  • Figure 7 describes the dynamic series of measurements with constantly changing glucose concentrations between 1 and 7.5 mM.
  • the individual measuring points (x) were determined from the slope of their potential transients between 10 and 20 ms after the charging pulse.
  • the squares ( ⁇ ) show independently performed amperometric measurements in correlation with the pulse potential transients.
  • the trace drawn was determined from the running average of 4 individual pulse measurements.
  • Figure 8 shows a glucose measurement on a graphite microelectrode with a diameter in the range of 350 pm and a length in the range of 1000 pm.
  • the coating is applied by sequentially immersing the graphite rod electrode in 1) 30 s in 1.25 mg/ml redox polymer, 2) 60 s pause, 3) 30 s in 2.5 mg/ml FAD-GDH, 4) 60 s pause, 5) 30 s in 1.25 mg/ml redox polymer, 6) drying for 1 h, 7) 15 min in 5 mg/ml crosslinker formed.
  • Figure 9 shows the measurement of a polymer film with a molecular weight ⁇ 5000 kD.
  • Figure 10 describes the measurements with an oxygen-independent FAD-dependent lactate dehydrogenase. Lactate is determined using LDH.
  • Figure 11 shows the compound or polymer Poly(MPC-GMA.
  • the microelectrodes can be coated with this compound or polymer Poly(MPC-GMA as a protective layer for the biocatalytic film. Coating with this compound or polymer is advantageous because it carries zwitterionic groups that prevent biofouling (protein and cell deposits).
  • Figure 12 finally discloses measurement of glucose with a poly(MPC-GMA) top layer.

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Abstract

L'invention se rapporte à un biocapteur implantable et à un procédé de mesure d'analyte automatisée continue dans un équilibre de concentration, comprenant une microélectrode et un film biologiquement actif disposé sur cette dernière et contenant un polymère redox qui adhère à la surface de la microélectrode au moyen d'un agent de réticulation, et une enzyme résistante à l'oxygène en tant que centre catalytique. Le biocapteur est conçu de telle sorte qu'un potentiel électrique, qui est plus négatif que le potentiel du polymère redox utilisé dans des conditions standard, est appliqué de manière permanente au biocapteur et seulement interrompu par la mesure d'analyte, l'application d'un potentiel à un intervalle de temps allant jusqu'à 2 secondes, c'est-à-dire le potentiel du polymère redox utilisé dans des conditions standard, déclenchant la mesure d'analyte et le changement potentiel dans le circuit à courant ouvert permettant la détermination de la concentration de l'analyte au niveau du biocapteur. L'invention se rapporte également à un procédé de mesure correspondant et à l'utilisation d'un capteur de ce type pour une implantation minimalement invasive dans le domaine de l'endocrinologie.
PCT/DE2023/100195 2022-03-15 2023-03-14 Biocapteur implantable WO2023174485A1 (fr)

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Citations (14)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
EP0778897A1 (fr) 1994-09-01 1997-06-18 Adam Heller Electrode sous-cutanee de surveillance du glucose
DE19814761A1 (de) 1998-04-02 1999-10-07 Merck Patent Gmbh Verfahren und Mittel zur Bestimmung von Proteinasen
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