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WO1997015827A9 - Capteur a fil enrobe - Google Patents

Capteur a fil enrobe

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Publication number
WO1997015827A9
WO1997015827A9 PCT/US1996/017893 US9617893W WO9715827A9 WO 1997015827 A9 WO1997015827 A9 WO 1997015827A9 US 9617893 W US9617893 W US 9617893W WO 9715827 A9 WO9715827 A9 WO 9715827A9
Authority
WO
WIPO (PCT)
Prior art keywords
sensor
enzyme
wire
platinum wire
platinum
Prior art date
Application number
PCT/US1996/017893
Other languages
English (en)
Other versions
WO1997015827A1 (fr
Filing date
Publication date
Application filed filed Critical
Publication of WO1997015827A1 publication Critical patent/WO1997015827A1/fr
Publication of WO1997015827A9 publication Critical patent/WO1997015827A9/fr

Links

Definitions

  • the present invention relates to a coated wire sensor, which can be used in the laboratory, has clinical applications, has applications in the food and pharmaceutical industries, and in particular can be used in the management of hemorrhagic shock.
  • a coated wire sensor can be used to increase the survival rate of injured patients who are at risk of hemorrhagic shock, especially while they are being transported to a hospital .
  • the important point in such an application is that unlike diabetes, hemorrhagic shock requires direct intravascular implantation of such sensors, and the monitoring of glucose levels for up to several hours.
  • hemorrhagic shock is characterized by the alteration of the glucose concentrations between the hypoglycemic and hyperglycemic levels.
  • a linearity of the sensor signal versus the glucose concentration is required. This linearity should be at least 22 mM (400 mg/dL) to assure linear output of the sensor within a safety margin.
  • the heretofore known devices do not provide for a high enough linearity.
  • Electropolymerization involves the electrochemical oxidation of a monomer from a solution containing the enzyme to form a conducting or non-conducting polymer on the electrode surface. This approach has several advantages. First, this process can be governed by the electrode potential, and therefore allows accurate control of the polymer film thickness and hence the amount of entrapped enzyme. Second, since the polymerization occurs locally at the electrode surface, it can be used to confine an enzyme precisely at the electrode without cross-immobilizing it on a neighboring electrode. This property is suitable for the fabrication of micro-arrays. Third, it is possible to use this technique for building multilayer structures, either one or more enzymes layered within a single polymer, or one enzyme within a multilayered copolymer.
  • Foulds et al also demonstrated the possible use of covalently attached ferrocenes to pyrrole monomers as a step towards the fabrication of reagentless sensors (N.C Foulds and CR. Lowe, Anal. Chem., 60 (1988) 2473) .
  • mediators in pyrrole films as anionic counter ions (Y. Kajiya, H. Sugai, C Iwakura and H. Yoneyama, Anal. Chem., 63 (1991 49; S. Yabuki, H. Shinohara, Y. Ikariyama and M. Aizawa, J. Electroanal . Chem. 277 (1990) 179) .
  • Non-conducting polymers have been studied as they have several advantageous characteristics not found in the conducting polymers. First, their growth is self limiting, thus producing thin polymer films. This results in a short sensor response time. The thin film also allows a higher enzyme content leading to a large sensor response and a high sensitivity. Non-conducting polymers also have a characteristic permselectivity, greatly decreasing the effect of common physiological interferents on the sensor response.
  • a non-conducting polymer that has been the focus of much research is 1, 2-phenylenediamine.
  • Sasso et al proposed a glucose sensor having a platinized reticulated vitreous carbon electrode with GOD entrapped in a poly(1, 2-phenylenediamine) film (S.V. Sasso, R.J. Pierce, R. Walla, and A.M. Yacynych, Anal. Chem., 62 (1990) 1111) . Their sensor was used in Flow Injection Analysis for determining glucose levels in human serum. They demonstrated the ability of the poly (1, 2-phenylenediamine) film in increasing the thermal stability of the GOD enzyme.
  • Malitesta et al constructed a glucose sensor, employing the same principle, but using platinum electrodes (C Malitesta, F. Palmisano, L.
  • poly(1, 3-phenylenediamine) Another non-conducting polymer that has been studied is poly(1, 3-phenylenediamine) .
  • This polymer is not commonly used in the construction of glucose sensors although it possesses the same p e r m s e l e c t i v e p r o p e r t i e s a s poly(1, 2-phenylenediamine) .
  • Geise et al . have demonstrated the use of 1, 1' -dimethylferrocene as a mediator entrapped with GOD in a poly(1, 3-phenylenediamine) film (R.J. Geise, S.Y. Roach and A.M. Yacynych, Anal. Chim.
  • Fig. 1 is a cross-sectional view of a first exemplary embodiment of the inventive coated wire sensor
  • Fig. 2 is a cross-sectional view of a second exemplary embodiment of an inventive coated wire sensor
  • Fig. 3 shows a third exemplary embodiment of an inventive coated wire sensor. Disclosure of the Invention
  • the coated wire sensor of the present invention is characterized primarily by: a platinum wire as a working electrode, a sensor body in the form of a hollow member disposed about at least a portion of the platinum wire, with the sensor body being spaced from the wire, a reference electrode that is operatively associated with a first end of the platinum wire but is not in contact with the wire, insulation disposed along at least a portion of the length of the platinum wire between the wire and the sensor body and/or the reference electrode, enzyme-containing material disposed on the first end of the platinum wire, with this enzyme-containing material comprising enzyme chemically linked to fine carbon particles, and a coating disposed over the enzyme- containing material.
  • the inventive electrode has a Pt base electrode surface.
  • Others use a carbon based electrode surface with a platinized surface - platinium hexachloride particles - a carbon ultramicroelectrode, using carbon fibers. We do not use a mediator. If a mediator is used, then a Pt surface is not needed.
  • the inventive electrode uses fine carbon particles plus Nafion, with the enzyme chemically bound (immobilized, linked to) the surface of the particles.
  • the inventive electrode has a very large effective surface area. This allows there to be an excess of enzyme in the pool, over that entrapped in the Nafion or the electrodeposited layer. Therefore, the inventive electrode function is independent of the degree of inactivation of the enzyme.
  • the inventive electrode works in a diffusion-limited mode, the activity of the electrode depends on the properties of the membranes employed, not on the degree of inactivation of the enzyme. Others use free enzyme above the base electrode, and sometimes an electron transfer mediator (such as ferrocene, resorcinol) . Construction of their electrodes suggests that their electrodes function in a kinetic regime.
  • the inventive electrode operates at an applied voltage (bias) of 0.65V. Others use a different bias, operating their electrode in a different electrochemical regime. 5.
  • the inventive electrode incorporates into the layer on top of the Pt base electrode, a Nafion ionomer to stabilize the sensor and decrease interference. Further, it makes the electrode function in a diffusion-limited mode, and also avoids many interference effects from external body chemicals or chemicals from an industrial process.
  • the inventive electrode has coatings applied on the outside of one or more of PVC, polyurethane, cellulose acetate, silastic This protects the internal enzymes from external interfering chemicals and fluids.
  • the inventive signal is, compared to others, very large, stable, and noise free.
  • the inventive sensor lifetime is up to 56 days, others achieve at best a few days.
  • the inventive sensor response time is a few seconds to a minute, others are 2 minutes or more.
  • the inventive sensors have a linear range up to 30 mM glucose in blood, others have only a linear range up to about 6-8 mM in buffers, with the use of toxic mediators. Other sensors cannot be used as implantable glucose sensors in humans, as they funtion and respond properly only over the normal to normal-high range of glucose concentration, not into the hyperglycemic range.
  • the inventive sensors responded in blood and plasma up to 38 mM glucose. Others had responses up to maybe 20 mM before the response flattens, and only with the use of toxic mediators (such as resorcinol, ferrocene etc) , undesirable in patients in shock.
  • toxic mediators such as resorcinol, ferrocene etc
  • Fig. 1 shows an inventive sensor that was fabricated by entrapping the glucose oxidase enzyme in an electrochemically grown poly(1,3- phenylenediamine) film.
  • This sensor which is, for example, a needle glucose biosensor, has a working electrode or anode in the form of the platinum wire 1. Except for the end 1' , the platinum wire 1 is covered by insulation 2, such as Silastic. Surrounding the insulation 2 is a sensor body 3, which in this embodiment is in the form of a stainless steel needle. Disposed on the exposed bulb end 1' of the platinum wire 1 is enzyme-containing material, in this embodiment in the form of enzyme incorporated in poly(1,3- phenylenediamine) film. A polymer coating 5 is then disposed over the enzyme-containing material 4. Finally, a reference electrode or cathode 6 is connected to the sensor body or needle 3.
  • a portion of a platinum wire working electrode IA is again covered with an insulation 2A, such as Silastic, leaving exposed one end of the wire 1.
  • the sensor body or stainless steel needle 3A is disposed on the insulation 2A and in this embodiment extends not only beyond the insulation but also beyond the exposed end l'A of the wire IA.
  • the enzyme-containing material 4A that is disposed over the end l'A is in this embodiment enzyme incorporated in a platinum black matrix.
  • a coating 5A for example, Nafion.
  • a reference electrode or cathode 6A is again connected to the needle 3A.
  • Fig. 3 shows a modified embodiment that can be used in the laboratory and for clinical applications.
  • the platinum wire IB has a portion thereof, except for the bulb-like end l'B, covered by insulation 2B in the form of a heat shrunk tube.
  • insulation 2B in the form of a heat shrunk tube.
  • a sensor body 3B Disposed about part of the insulation 2B is a sensor body 3B in the form of an outer wrapping that is preferably made of plastic, although it could also be made of some other material, such as metal.
  • a reference electrode or cathode 6B extends between the insulation 2B and the sensor body 3B and is wrapped around the insulation 2B between the bulb ⁇ like end l'B of the platinum wire and the sensor body 3B.
  • the enzyme- containing material 4B comprises a conducting or non-conducting polymer.
  • Conducting and nonconducting polymers that can be used include, but are not limited to conducting polymers such as poly(pyrrole) , poly (N-methylpyrrole) (P.N. Bartlett and P.G. Whittaker, J. Electroanal. Chem., 224 (1987) 37) , poly(aniline) (J.C Cooper and E.A.H. Hall, Biosensors, 7 (1992) 473) ; nonconduct ing polymers such as poly ( l , 2 -phenylenediamine ) and poly(1, 3-phenylenediamine) , poly- (phenol) (J. Wang, S-P Chen and M-S Lin, J. Electroanal. Chem., 273 (1989) 231; R.L.
  • conducting polymers such as poly(pyrrole) , poly (N-methylpyrrole) (P.N. Bartlett and P.G. Whittaker, J. Electroanal. Chem., 224 (1987) 37) , poly
  • Polymer membranes that can be used for the coating 5, as well as the insulation 2 include, for example, polyurethane, polyvinylchloride, Silastic, and cellulose acetate, polytetrafluoroethylene, polycarbonate.
  • the design approach for fabrication of the sensor was chosen to be that of miniaturization to a needle size.
  • a hypodermic needle approximately 18 gauge -- outer diameter 1.27 mm was chosen to act as sensor body and as counter and reference electrodes .
  • three different methods of enzyme immobilization were chosen and investigated, leading to 3 different prototypes.
  • the fabrication of the sensors involved three principal steps.
  • the first step which is common to the three sensor designs, involves the initial pretreatment and preparation of the needle body and the platinum electrode.
  • the second step is the immobilization of the glucose oxidase enzyme and its incorporation into the needle body.
  • the third and last step is the formation of the diffusion limiting membrane or coating. The coatings differ according to each immobilization technique.
  • the initial preparation and pretreatment of the sensor body and platinum wires were as follows.
  • the stainless steel needle was cut at both ends to remove the plastic cap and the pointed end, and the ends were smoothed using files and fine sand paper.
  • the needle was cleaned in concentrated nitric acid, washed with distilled water and dried.
  • a platinum wire was cleaned in concentrated nitric acid and treated in a propane flame to form a smooth shape or a bulb at one of the ends.
  • the platinum wire was insulated by dipping in Silastic up to 2 mm below the bulb end, or by sealing in a polyethylene tubing with proper inner diameter and was fixed in place by gluing to the sensor body with cyanoacrylate glue. Wires of 0.5 mm diameter were insulated using Teflon heat shrinkable tubing.
  • the partially insulated platinum wire was electrochemically treated by switching between +2.0 V and -1.0 V versus a silver/silver chloride electrode in a three electrode scheme six times for one hour.
  • Figure 1 shows a schematic of a sensor fabricated by entrapping the GOD enzyme in an electrochemicallygrownpoly(1, 3-phenylenediamine) film.
  • the electropolymerization of 1, 3-phenylenediamine and the incorporation of the enzyme particles was carried out in a solution containing 3-5 mg of 1, 3-phenylenediamine, 20 mg of Glucose Oxidase and 20 mg of ULTI carbon powder in 9 mL at pH 7.4 and 1 mL of Nafion.
  • This preparation process was performed potentiostatically at +0.65 V versus a silver/silver chloride reference electrode for 15 minutes.
  • An additional coating of poly(1, 3-phenylenediamine) was carried out. The formation of this coating was in a solution of 3-5 mg of 1, 3-phenylenediamine .
  • This process was performed potentiostatically at +0.65 V versus a silver/silver chloride reference electrode for 10 minutes.
  • the prepared platinum electrode was washed thoroughly in a stirred buffer solution overnight .
  • a variety of different polymer membranes of varying concentrations were applied to this sensor. The membranes were formed over the sensor tip by dip casting. The sensor end was dipped in the polymer solution for 20 seconds. The sensor was then held vertically and dried in air for one hour. The results of evaluation tests on the above manufactured sensors are given below.
  • Figure 2 shows a schematic of a sensor fabricated by entrapping the GOD enzyme in a platinum black matrix.
  • the platinization and electrophoretic incorporation of the enzyme particles was carried out in a solution containing 33 mg of sodium hexachloroplatinate and 30 mg of Glucose Oxidase and 0.6 mg of lead acetate in one mL at pH 3.5 (Y.
  • Nafion coatings of the sensor were obtained by dipping the sensor end in 0.5% ionomer solution for 30 seconds. The sensor was then held vertically and dried in air for one hour.
  • Fig. 3 shows the schematic of the electrode construction.
  • the electrode consists of a platinum wire, a coating layer of St ⁇ ber glass beads on the platinum, and the outer immobilized enzyme layer.
  • the glass coating layer has two effects: it acts as a support and matrix for the enzyme; it can be a barrier to some substances that may cause disturbance in the signal.
  • Silanization was utilized to prepare the St ⁇ ber glass beads and the platinum wire surfaces to bond with glutaraldehyde.
  • a solution of 3- aminopropytriethoxy silane (0.2 mL) in distilled water (2 mL, to give a 10% v/v solution) was prepared.
  • the pH was adjusted to 4.1 using hydrochloric acid (HCl) solution.
  • This mixture was subsequently placed into a small vial and several electrode immersed into the vial . They were kept in an oven at 80°C and shaken every fifteen minutes for three hours.
  • the sensors were rinsed completely in distilled water and then dried at 120°C for three hours.
  • the wires were stored in the refrigerator for 24 hours.
  • Immobilization was accomplished by first creating a layer of glutaraldehyde bonded to the silanized electrode surface.
  • a 2.5% glutaraldehyde solution was made with 27 mL distilled water and 3 mL of 25% glutaraldehyde.
  • Immobilization was accomplished by immersing the wire in the glutaraldehyde solution, continuously stirred, for one hour at room temperature. The wires were then rinsed with water for one hour at room temperature, after which the electrodes were immersed in the enzyme solution.
  • the enzyme solution was prepared with 80 mg of glucose oxidase in 2 mL phosphate buffer of pH 7.4. The wires were left in the enzyme solution overnight at 4°C to allow immobilization onto the glass layer.
  • Table 1 shows a summary of the sensor characteristics obtained with sensors employing five different polymer coatings, namely; cellulose acetate (CA) , Silastic 3-5025, Silastic Q7-2213, polyurethane (PU) and polyvinylchloride (PVC) .
  • Table 1 presents a variety of concentrations of the polymer coating solutions and the resulting parameters of the calibration curves of the sensors: linearity (as the upper limit of the calibration curve linear range) , sensitivity (slope of the linear portion of the calibration curve) and response time. It should be noted that any coated sensors showing a non-linear response were assumed to have a linear range of 2.2 mM as the linearity below this range was not investigated.
  • Table 2 shows the evaluation of the second prototype sensor performance in plasma and in buffer solution before and after the plasma test. Comparing the sensor response in buffer solution before the plasma test with that during the plasma test, it can be seen that the values of the sensor response to glucose concentration from 10 to 12 mM in both measurements coincide (within 2%) . At higher glucose concentration the signal in plasma is lower than that obtained in buffer solution but with an acceptable deviation of 7%. A higher response of the sensor to glucose levels in plasma within the range 4.4 - 8.4 mM is observed. The values of the sensor response obtained in buffer after the plasma test practically coincide with those obtained in plasma.
  • the effect of interferents was tested by five common substances.
  • the first sensor prototype shows a decrease in sensitivity after the addition of interferents.
  • the second sensor prototype shows a decreased effect of glycine, urea and ascorbic
  • Hemorrhagic shock is characterized by the alteration of glucose concentrations between the hypoglycemic and hyperglycemic levels.
  • the sensor response should be linear in glucose concentration over the range of interest. This linearity should be at least 22 mM (400 mg/dL) to assure linear output of the sensor within some safety margin. It can be seen from Table 4 that all sensor prototypes satisfy this requirement . The linearity of the sensor output signal versus the glucose concentration is extended to at least 22 mM.
  • the first sensor prototype showed a linearity of 26.7 mM with sensitivity of 1.62 nA/mM, 31.1 mM with a sensitivity of 1.35 nA/mM, and 37.7 mM with a sensitivity of 1.8 nA/mM for cellulose acetate (CA) , polyurethane (PU) and polyvinylchloride (PVC) coated sensors, respectively.
  • the second sensor prototype (with enzyme entrapped in a platinum black matrix) showed a linearity of up to 33 mM.
  • the third sensor prototype also showed a linearity of at least 22 mM. Another requirement for monitoring glucose levels during hemorrhagic shock is a fast sensor response (short response time) to changes in glucose concentration.
  • the first sensor prototype shows response times of 183, 24 and 35 seconds for CA, PU and PVC coated sensors, 5 while the second sensor prototype shows a response time of 60 seconds; the third prototype shows a resonse time of 90 seconds.
  • Another test reflecting the response times of the sensors is the in vitro hemorrhagic shock simulation test.
  • the second sensor prototype shows a response time of 5 minutes for the increasing and 10 minutes for the decreasing step changes. It is
  • the first and second sensor prototypes show the shortest response times in response to both small and large changes in glucose concentration.
  • Plasma Test Decreased Decreased No change in sensit ivity Sensit ivity Sensitivity
  • the reproducibility of the sensor signal 35 is good in all three sensor prototypes.
  • the variations in the sensor response do not exceed 5% in the first prototype response, while it is less than 10% for the second prototype.
  • All three prototypes showed a life time 40 of at least one week with no significant change in the sensor performance, hence satisfying the life time requirement .
  • One last aspect to be considered is the ease of fabrication of the sensor prototypes. In general, the three sensor prototypes are easily fabricated. The first prototype has the advantage of electrochemical growth of the entrapping film, hence allowing better controllability on the thickness of the film and hence the amount of enzyme entrapped.

Abstract

Ce capteur à fil enrobé comprend un fil de platine (1) en guise d'électrode opérationnelle, un corps (3) de capteur, creux et disposé autour d'au moins une partie du fil de platine, fil dont il est séparé, une électrode de référence (6) associée opérationnellment avec une première extrémité du fil de platine mais sans entrer en contact avec ce fil, une isolation (2) disposée au moins le long d'une partie du fil de platine entre celui-ci et le corps de capteur et/ou l'électrode de référence, une substance (4) contenant une enzyme, disposée à la première extrémité du fil de platine et contenant une enzyme liée chimiquement à de fines particules de carbone, ainsi qu'un enrobage (5) disposé sur cette substance contenant une enzyme.
PCT/US1996/017893 1995-10-25 1996-10-25 Capteur a fil enrobe WO1997015827A1 (fr)

Applications Claiming Priority (2)

Application Number Priority Date Filing Date Title
US54781495A 1995-10-25 1995-10-25
US08/547,814 1995-10-25

Publications (2)

Publication Number Publication Date
WO1997015827A1 WO1997015827A1 (fr) 1997-05-01
WO1997015827A9 true WO1997015827A9 (fr) 1997-11-13

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JP3398598B2 (ja) * 1998-06-10 2003-04-21 松下電器産業株式会社 基質の定量法ならびにそれに用いる分析素子および測定器
AU1631101A (en) * 1999-10-07 2001-05-10 Pepex Biomedical, Llc Sensor for measuring a bioanalyte such as lactate
DE10015818A1 (de) 2000-03-30 2001-10-18 Infineon Technologies Ag Biosensor und Verfahren zum Ermitteln makromolekularer Biopolymere mit einem Biosensor
DE10015816A1 (de) * 2000-03-30 2001-10-18 Infineon Technologies Ag Biosensorchip
DE10032042A1 (de) * 2000-07-05 2002-01-24 Inventus Biotec Gesellschaft Fuer Innovative Bioanalytik, Biosensoren Und Diagnostika Mbh & Co. Kg Elektrochemischer Einwegbiosensor für die quantitative Bestimmung von Analytkonzentrationen in Flüssigkeiten
DE10119036C1 (de) 2001-04-18 2002-12-12 Disetronic Licensing Ag Tauchsensor zur Messung der Konzentration eines Analyten mit Hilfe einer Oxidase
US6872297B2 (en) 2001-05-31 2005-03-29 Instrumentation Laboratory Company Analytical instruments, biosensors and methods thereof
US6652720B1 (en) * 2001-05-31 2003-11-25 Instrumentation Laboratory Company Analytical instruments, biosensors and methods thereof
AU2006252048B2 (en) * 2001-05-31 2008-09-04 Instrumentation Laboratory Company Analytical instruments and biosensors, and methods for increasing their accuracy and effective life
US6960466B2 (en) 2001-05-31 2005-11-01 Instrumentation Laboratory Company Composite membrane containing a cross-linked enzyme matrix for a biosensor
US6501976B1 (en) * 2001-06-12 2002-12-31 Lifescan, Inc. Percutaneous biological fluid sampling and analyte measurement devices and methods
WO2009051901A2 (fr) 2007-08-30 2009-04-23 Pepex Biomedical, Llc Capteur électrochimique et procédé de fabrication
WO2009032760A2 (fr) 2007-08-30 2009-03-12 Pepex Biomedical Llc Capteur électrochimique et procédé de fabrication
US8951377B2 (en) 2008-11-14 2015-02-10 Pepex Biomedical, Inc. Manufacturing electrochemical sensor module
US9445755B2 (en) 2008-11-14 2016-09-20 Pepex Biomedical, Llc Electrochemical sensor module
US9585605B2 (en) 2011-05-19 2017-03-07 Pepex Biomedical, Inc. Fluid management and patient monitoring system
WO2012162151A2 (fr) 2011-05-20 2012-11-29 Pepex Biomedical, Inc. Fabrication de modules de détection électrochimique
CN104918551B (zh) 2012-12-03 2019-07-26 Pepex生物医药有限公司 传感器模块以及使用传感器模块的方法
US9594047B2 (en) * 2013-04-26 2017-03-14 Universiteit Antwerpen Potentiometric sensors and method for measuring intermolecular interactions
CN107003264B (zh) 2014-06-04 2020-02-21 普佩克斯生物医药有限公司 电化学传感器和使用先进印刷技术制造电化学传感器的方法

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DE3700119A1 (de) * 1987-01-03 1988-07-14 Inst Diabetestechnologie Gemei Implantierbarer elektrochemischer sensor
FR2630546B1 (fr) * 1988-04-20 1993-07-30 Centre Nat Rech Scient Electrode enzymatique et son procede de preparation
US5089112A (en) * 1989-03-20 1992-02-18 Associated Universities, Inc. Electrochemical biosensor based on immobilized enzymes and redox polymers

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