Anal. Chem. 1994,66, 3131-3138
Design, Characterization, and One-Point in Vivo Calibration of a Subcutaneously Implanted Glucose Electrode Eiisabeth Csoreg1,t Chris P. Quinn,* David W. Schmldtke,* Sten-Eric Lindquist,g Michael V. Pishko,* Ling Ye,* Ioanis Katakis,* Jeffrey A. Hubbell,* and Adam Heiier’** Department of Chemical Engineering, University of Texas at Austin, Austin, Texas, 787 12- 1062, Department of Analytical Chemistry, University of Lund, P.O. Box, 124, S-221 00 Lund, Sweden, and Department of Physical Chemistry, University of Uppsala, P.O. Box 532, S-751 21 Uppsala, Sweden
A 0.29-mm-diameter flexible electrode designed for subcutaneous in vivo amperometric monitoring of glucose is described. The electrode was designed to allow “one-point” in vivo calibration, Le., to have zero output current at zero glucose concentration, even in the presence of other electroreactive species of serum or blood. A valid zero point, along with a measurement of the glucose concentration in a withdrawn sample of blood at which the current is known, defined the sensitivity in the linear response range. The electrode was four-layered, with the layers serially deposited within a 0.125mm recess upon the tip of a polyimide-insulated 0.25-mm gold wire. The recessed structure reduced the sensitivity to movement and allowed, through control of the depth of the recess, control of the transport of glucose and thus the range of linearity. The recess contained the four polymeric layers, with a total mass less than 5 pg and no leachable components. The bottom glucose concentration-to-current transducing layer consisted of the enzyme “wiring” redox polymer poly[(~inylimidazole)Os(bipyridine)2Cl]+~~+, complexed with recombinant glucose oxidase and cross-linked with poly(ethy1ene glycol) diglycidyl ether, to form an electron-conducting hydrogel. The layer was overcoated with an electrically insulating layer of polyaziridine-cross-linked poly (allylamine), on which an immobilized interference-eliminating horseradish peroxidase based film was deposited. An outer biocompatible layer was formed by photo-cross-linking derivatized poly(ethylene oxide). The current output of a typical electrode at 10 mM glucose and at 37 OC was 35 nA, the apparent K, was 20 mM, and the 10-90% response time was -1 min. The sensitivity varied only by f5% over a 7 2 4 test period. The electrode tracked the blood glucose concentration in a rat model before, during, and following intraperitoneal glucose infusion. Two failure modes were observed. The first, deactivation of hydrogen peroxide-producing lactate oxidase in the interference-eliminating layer, resulted in inadequate preoxidation of interferants. The second was an abrupt drop in the sensitivity of implanted electrodes, -7 h after implantation. The failed electrodes promptly regained their sensitivity in buffer. In response to the need for frequent or continuous in vivo monitoring of glucose in diabetics, particularly in brittle diabetics, a range of possible in vivo glucose electrodes have been Thedesired characteristics of theseelectrodes + University of Lund. t
University of Texas at Austin.
I University of Uppsala.
0003-2700/94/0366-3131$04.50/0 0 1994 American Chemical Society
include (a) safety, (b) clinical accuracy and reliability, (c) feasibility of in vivo recalibration, (d) stability for at least one hospital shift of 8 h, (e) small size, ( f ) ease of insertion and removal, and (g) a sufficiently fast response to allow timely intervention. Keys to safety are absence of leachable components, biocompatibility, and limiting of the potentially hazardous foreign matter introduced into the body to an amount that is inconsequential in a worst case failure. The clinical accuracy (1) Armour, J. C.; Lucisano, J. Y.; McKean, B. D.; Gough, D. A. Diabetes 1990, 39, 1519-1526. (2)Shichiri, M.; Kawamori, R.; Yamasaki, Y. Merhods Enzymol. 1988, 137, 326-334. (3)Shichiri, M.; Yamasaki, Y.; Nao, K.; Sekiya, M.; Ueda, N. In Zmplanrable Glucose Sensors-The Srare of rhe Art; Pfeiffer, E. F., Reaven, G. M., Eds.; Supplement Series/Hormone and Metabolic Research; (Reisensburg, 1987) Thieme Medical Publishers, Inc.: New York, 1988;pp 17-20. (4) Moatti-Sirat, D.;Capron, F.; Poitout, V.; Reach, G.; Bindra, D. S.; Zhang, Y.; Wilson, G. S.; Thevenot, D. R. Diaberologia 1992,35,224230. (5) Karube, 1.; Yokoyama, K.; Tamiya, E. Biosens. Bioelecrron. 1993,8, 219228. (6) Brilckel, J.; Kerner, W.; Zier, H.;Steinbach, G.; Pfeiffer, R. F. Klin. Wochenschr. 1989,67,491495. (7)Gernet, S.; Koudelka, M.; Rwij de, N. F. S e w . Acruarors 1989,17,537-540. (8)Pickup, J. C.; Claremont, D. J.; In Zmplanrable Glucose Sensors-The Srare ojthe Art;Pfeiffer, E. F., Reaven, G. M., Eds.;Supplement Series/Hormone and Metabolic Research; (Reisensburg, 1987)Thieme Medical Publishers, Inc.: New York, 1988;pp 3636. (9)Yokoyama, K.; Lee, S. M.; Tamiya, E.; Karube, I.; Nakajima, K.; Uchiyama, S.; Suzuki, S.; Akiyama M.; Masuda, Y. Anal. Chim. Acra 1992,263,101110. (10)Tamiya, E.; Karube, I. Sens. Acruarors 1988,15, 199-207. (11) Bindra, D.S.; Zhang, Y.; Wilson, G. S.; Sternberg, R.; Thevenot, D. R.; Moatti, D.; Reach, G. Anal. Chem. 1991,63,1692-1696. (12)Cronenberg, C.; Grocn van, B.; Beer de, D.; Heuvel van den, H. Anal. Chim. Acta 1991, 242,275-278. (13) Johnson, K. W.; Mastrototaro, J. J.; Howey, D. C.; Brunella, R. L.; Burden Brady, P. L.; Bryan, N. A.; Andrew, C. C.; Rowe, H. M.; Allen, D. J.; Noffke, B. W.; McMahan, W. C.; Morff, R. J.; Lipson, D.; Nevin, R. S . Biosens. Bioelectron. 1992,7, 709-714. (14)Koudelka, M.; Gernet, S.;Rwij de, N. F. Sens. Acruarors 1989,18,157-165. (15)Shaw, G. W.; Claremont, D. J.; Pickup, J. C. Biosens. Bioelectron. 1991, 6, 401406. (16)Pickup, J. C.; Shaw, G. W.; Claremont, D. J. Diaberologia 1989,32,213-217. (17)Gregg, B.; Heller, A. Anal. Chem. 1990,62,258-263. (18) Gregg, 8.; Heller, A. J. Phys. Chem. 1991, 95,5976-5980. (19)A h , T.; Lau, Y. Y.; Ewing, A. G. AMI. Chem. 1992,64,2160-2163. (20)Bartlett, P. N.; Caruana, D. J. Analyst 1992,117, 1287-1292. (21)Chen, C.-Y.; Gotoh, M.; Makino, H.;Su, Y.-C.; Tamiya, E.; Karube, I. Anal. Chim. Acra 1992,265, 5-14. (22)Kawagoe, J. L.; Niehaus, D. E.; Wightman, R. M. Anal. Chem. 1991, 63, 2961-2965. (23)Murakami, T.; Nakamoto, S.; Kimura, J.; Kuriyama, T.; Karube, I. Anal. Lett. 1986,19,1973-1986. (24)Pishko, M. V.;Michael, A. C.; Heller, A. Anal. Chem. 1991,63,2268-2272. (25)Tamiya, E.;Sugiura, Y.; Akiyama, A.; Karube, I. Ann. N.Y. Acad. Sci. 1990, 613. 396-400. (26)Turner, R. F. B.; Harrison, D. J.; Rajotte, R. V.; Bakes, H. P. Sens. Acruarors B 1990,21, 561-564. (27)Urban, G.; Jobst, G.; Keplinger, F.; Aschauer, E.; Tilado, 0.; Fasching, R.; Kohl, F. Biosens. Bioelectron. 1992,7,733-739. (28)Velho, G.; Froguel, P.; ThCenot, D. R.; Reach, G. Diabetes Nurr. Merab. 1988,I , 227-234.
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must be such that even when the readings are least accurate, the clinical decisions based on these be still correct. Feasibility of prompt confirmation of proper functioning of the sensors and of periodic in vivo recalibration is of the essence if a physician is to allow the life of a patient to depend on the sensor. This one-point calibration, relying on the signal at zero glucose concentration being zero and measuring the blood glucose concentration at one point in time, along with the signal, is essential, but has heretofore been elusive. The sensitivity must be sufficiently stable for the frequency of the required in vivo recalibration to not be excessive. The sensor must be small enough to be introduced and removed with minimal discomfort to the patient and for minimal tissue damage. It is preferred that the sensor be subcutaneous and that it be inserted and removed by the patient or by staff in a physician’s office. Finally, its response time must be fast enough so that the corrective measures, when needed, will be timely. In response to some of these needs, needle typel.2A6.8.11-13,16,19,22,23,25-27,29-32and other subcutaneous amperometric sensors were considered. The majority of these utilized platinum, platinum-iridium, or platinum black to electrooxidize H202 generated by the glucose oxidase (GOx) catalyzed reaction of glucose and oxygen.2,5q6J1-13~16~19~26,27*30 In these sensors, the GOx was usually in large excess and immobilized, often by cross-linking with albumin and g l ~ t a r a l d e h y d e . ~ ~ To - ~ ~exclude ? ~ ~ , ~electrooxidizable ~ interferants, membranes of cellulose acetate and sulfonated polymers including Nafion were ~sed.~~,~6,33,34 Particular attention was paid to the exclusion of the most common electrooxidizable interferants: ascorbate, urate, and acetaminophen. Also, to cope with the interferants, two-electrode differential measurements were used, one electrode being sensitive to glucose and electrooxidizable interferants and the other only to i n t e r f e r a n t ~ .A~ ~strategy for overcoming the problem of interferants, applied also in this work, involved their p r e ~ x i d a t i o n . ~To ~ . ~make ~ the electrodes more biocompatible, hydrophilic polyurethanes,2.6*11.37,38 poly(viny1 a l ~ o h o l ) and , ~ ~ o ~ ~ H Emembranes M A ~ ~have been used. Several researchers tested GOx-based glucose sensors in vivo and obtained acceptable results in rats,4*30,37*42 rabbit~,13?~3 d 0 g ~ , 2 , pigs,8-41142 ~ + ~ ~ , ~sheep,6and h u m a n ~ . l ~These , ~ ~ studies >~ validated the subcutaneous tissue as an acceptable glucose(29) Wang, J.; Angnes, L. Anal. Chem. 1992,64, 456-459. (30) Koudelka, M.; Rohner, J. F.; Terrettaz, J.; Bobbioni, H . E.; Rooij de, N . F.; Jeanrenaud, B. Biosens. Bioelectron. 1991, 6, 31-36. (31) Velho, G.; Froguel, P.; Sternberg, R.; Thevenot, D. R.; Reach, G . Diabetes 1989, 38, 164-171. (32) Pishko, M. V.; Katakis, I.; Lindquist, S.-E.; Heller, A. Mol. Cryst. Liq. Crysr. 1990, 190, 221-249. (33) Shiono, S.;Hanazato, Y.; Nakako, M. Anal. Sei. 1986, 2, 517-521. (34) Gorton, L.; Karan, H. I.; Hale, P. D.; Inagaki, T.; Okamoto, Y.; Skotheim, T.A. Anal. Chim. Acta 1990, 28, 23-30. (35) Maidan, R.; Heller, A. Anal. Chem. 1992, 64, 2889-2896. (36) Maidan, R.; Heller, A. J . Am. Chem. SOC.1991, 113, 9003-9004. (37) Koudelka, M.; Rohner, J. F.; Terrettaz, J.; Bobbioni, H. E.; Rooij de, N . F.; Jeanrenaud, B. Biomed. Biochim. Acta 1989.48.953-956. (38) Velho, G.; Froguel, P.; Thevenot, D. R.; Reach, G. Biomed. Biochim. Acta 1989, 48, 951-964. (39) Rebrin, K.; Fischer, U.; Woedtke, T.; Abel, P.; Brunstein, E. Diabetologia 1989, 32, 573-576. (40) Fischer, U.; Ertle, R.; Abel, P.; Rebrin, K.; Brunstein, E.; Dorschevon, H. H.; Freyse. E. J. Diabetologia 1987, 30, 940-945. (41) A h , T.; Lau, Y. Y.; Ewing, A. G. J. Am. Chem. Soc. 1991,113,7421-7423. (42) Claremont. D. J.; Pickup, J. C.; Sambrook, I. E. Life Support Sysr. 1986, 369-370. (43) Ito, K.; Ikcda, S.;Asai, K.; Naruse, H.; Ohkura, K.; Ichihashi, H.; Kamei, H.; Kondo, T. ACS Symp. Ser. 1986, 309, 373-382.
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sensing site. Good correlation was observed between intravascular and subcutaneous glucose concentration^.^^,^^ They also demonstrated the need for in vivo sensor ~ a l i b r a t i o n . ~ ~ Another approach to in vivo glucose monitoring was based on coupling subcutaneous microdialysis with electrochemical detection.46 To control and adjust the linear response range, electrodes have been made glucose-diffusion limited, usually through glucose transport limiting membranes. Diffusional mediators, through which the 02 partial pressure dependence of the signals is reduced, are leached from sensors. Such leaching introduces an unwanted chemical into the body and also leads to loss in sensitivity, particularly in small sensors. In microsensors, in which outward diffusion of the mediator is radia1,47.48the decline in sensitivity is rapid. This problem has been overcome in “wired” enzyme electrodes, i.e., electrodes made by connecting enzymes to electrodes through cross-linked electron-conducting redox hydrogels Glucose oxidase has been wired with polyelectrolytes having electron relaying [Os(bpy)2C1]+/2+ redox centers in their backbones. Hydrogels were formed upon cross-linking the enzyme and its wire on electrodes. These electrodes had high current densities and operated at a potential of 0.3 V vs SCE.17*50The electrooxidizable interferants were eliminated through peroxidase-catalyzed preoxidation in a second, nonwired, hydrogen peroxide generating layer on the wired enzyme ele~trode.~5.36 In this paper, we report on a 0.29-mm recessed gold wire electrode for subcutaneous in vivo glucose monitoring, approaching in its performance all of the above listed requirements, including in vivo one-point calibration. The electrode was constructed by depositing four layers into a recess. The recess was formed by etching away gold from an insulated gold wire. The polymer layers were protected within the recess against mechanical damage. The recess and its polymer layers also reduced the transport of glucose to the gold contacting sensing layer. By limiting the glucose flux, the desired linear response range, spanning the clinically relevant glucose concentration range, was obtained. The sensor had no leachable components, and its four cross-linked polymer layers contained only 5 pg of immobilized material and only a few nanograms of polymer-bound osmium. Preoxidation of the interferants in one of the four layers made possible one-point in vivo calibration of the sensor.
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EXPERIMENTAL SECTION Materials. The electrodes were made of a polyimideinsulated 250-pm-diameter gold wire, 290-pm outer diameter (0.d.) (California Fine Wire Co., Grover City, CA). Heatshrinkable tubing (RNF 100 3/64 in. BK and ‘ / I 6 in. BK, Thermofit, Raychem, Menlo Park, CA) and a two-component (44) Shichiri, M.; Asakawa, N.; Yamasaki, Y.; Kawamori, R.; Abe, H. Diabetes Care 1986, 9, 298-301. (45) Kerner, W.; Zier. H.; Steinbach, G.; Briickel, J.; Pfeiffer, E. F.; Wiess, T.; Camman, K.; Plack, H. In Implantable Glucose Sensot-The State of the Art; Pfeiffer, E. M., Raven, G. M., Eds.; Supplement Serics/Hormone and Metabolic Rcaearch; (Reiscnsburg, 1987) Thieme Medical Publishers Inc.: New York, 1988; pp 8-13. (46) Moscone, D.; Pasini, M.; Mascini, M. Talanta 1992, 39, 1039-1044. (47) Aoki, K. Electroanalysis 1993, 5, 627639. (48) Fleischmann, M.; Pons, S.;Rolison, D. R.;Schmidt, P. P. Ultramicroelectrodes; Publishers Press: Morganton, NC, 1987. (49) Heller, A. J . Phys. Chem. 1992, 96, 3579-3587. (50) Ohara, T.; Rajagopalan, R.; Heller, A. Anal. Chem. 1993, 65, 24-29.
silver epoxy (Epo-tek H20E; Epoxy Tech. Inc., Billerica, MA) were used for electrode preparation. The glucose-sensing layer was made by cross-linking a genetically engineered glucose oxidase (rGOx) (35% purity, Chiron Corp., Emeryville, CA) with a polymer derived of poly(vinylimidazo1e) (PVI), made by complexing part of the imidazoles to [O~(bpy)2Cl]+/~+. The resulting redox polymer, termed PVI-Os, was synthesized according to a previously published protocol.50 Poly(ethy1ene glycol) diglycidyl ether 400 (PEGDGE; Polysciences, Warrington, PA) was used as the cross-linker. The barrier layer between the sensing and interferenceeliminating layers was made of poly(ally1amine) (PAL; Polysciences) cross-linked with a polyfunctional aziridine (PAZ) (XAMA-7; Virginia Chemicals, Portsmouth, VA). The interference-eliminating layer was prepared by coimmobilizing horseradishperoxidase (HRP) type VI (Catalog no. P-8375,3 10 units/mg, denoted herein as HRP-VI; Sigma, St. Louis, MO) and HRP for immunological assay (No. 8 14407, minimum 1000 units/mg, denoted HRP-BM; Boehringer-Mannheim, Indianapolis, IN) with lactate oxidase from Pediococcus sp. (Catalog No. 1361, 40 units/mg, denoted LOX; Genzyme, Cambridge, MA) and a recombinant microbial source (Catalog No. 138l denoted rLOx, Genzyme). Coimmobilization was performed using sodium periodate (Catalog No. S-1147, Sigma).35 The biocompatible layer was made of 10% aqueous poly(ethylene oxide) tetraacrylate (PEO-TA). To form the photocross-linkable polymer, PEO was acrylated by reaction with acryloyl chloride. The 18 500 g/mol PEO (Polysciences) is a tetrahydroxylated compound by virtue of two hydroxyl groups on a bisphenol A bisepoxide that linked two a, o-hydroxy-terminated 9000 g/mol PEO units. Acryloyl chloride (Aldrich, Milwaukee, WI) in a 2-5 M excess was used to acrylate the polymer (10% w/v PEO in benzene). Triethylamine (Mallinkrodt, Paris, KY) was used as a proton acceptor equimolar with the acryloyl chloride.51 Other chemicals used were ascorbic acid, uric acid, 4-acetaminophenol,L-(+)-lactic acid, and hydrogen peroxide (30%), all from Sigma. All chemicals were used as received. Solutions(if not otherwisespecified) were made with distilled, deionized water. Glucose monitoring was performed in buffer, in bovine serum (Sigma, Catalog No. S-6648) containing antibiotic-antimycotic solution (Sigma, Catalog No. A-8909) at 37 "C, and in rats. Instrumentation. In making the recessed gold electrodes, a potentiostat-galvanostat (PAR Model 173, Princeton Applied Research, Princeton,NJ), operated in a galvanostatic mode, and a sonicator (Fisher Scientific, Pittsburgh, PA) were used. Cyclic voltammograms were recorded with a potentiostat (PAR Model 273A) and a conventional electrochemical cell having a Pt wire counter and a SCE reference electrode and were evaluated with PAR 270 software. Glucose signals were monitored with a bipotentiostat (Biometra, EP 30, Gothingen, Germany) and a two channel strip-chart recorder (Kipp & Zonen, Bohemia, NY). The recessed electrodes were coated under a microscope (Bausch & Lomb) using a micromanipulator (Narishige,Seacliff, NY). The micropipets (51) Pathak, C. P.; Sawhney, A. 83 11-83 12.
S.;Hubbell, J. A. J. Am. Chem. Soc. 1993,115,
I
Poly imide
1. Sensing Layer
2. Insulating Layer
3. Interference Eliminating Layer 4. Biocompatibie Layer
Flgure 1. (A) Photomicrograph of a recessed, uncoated electrode, lying flat on a surface. Use of a light microscope afforded a view through the polyimide insulation but not through the gold wire. (B) Schematic drawing of the glucose sensor.
were pulled with a micropipet puller (Narishige). Temperature was controlled with an isothermal circulator (Fisher Scientific). Electrode Preparation. Lengths (5 cm) of polyimideinsulated gold wire were cut with a sharp razor blade. Electrical contact was made at one end with silver epoxy to an insulated stainless steel wire, and the junction was covered with insulating heat-shrinkable tubing. The recess-forming electrochemical etching process was carried out in 10 mL of 3 M potassium cyanide, with the gold wire as the working electrode and a platinum or gold wire as the counter electrode. The wires were placed in contact with the bottom of the beaker, all electrodes being equidistant from the counter electrode. The beaker was sonicated during the etching procedure. The ends of the Au wires were bent upward, so that agitation by the sonicator caused the oxygen bubbles formed during the etching process to rise and escape. The electrodes were then thoroughly washed and immersed in water for 30 min. The recession procedure is highly reproducible; a deviation off 10 pm was found (using an objective micrometer) for a batch of 30 recessed electrodes. Before coating, the electrodes were examined under a microscope for flatness of the gold surface and correct depth. Figure 1A shows a side view of a recessed, uncoated electrode. The recessed gold surfaces were coated by filling of the cavities with aqueous solutions containing the cross-linkable components of the different layers, and their cross-linkers. The solutions were introduced under a microscope with a micropipet (connected to a microsyringe by polyethylene tubing and shrink tubing), using a micromanipulator. After application of each of the individual layers, the electrodes were cured overnight at room temperature in air. Analytical Chemistry, Vol. 66, No. 19, October 1, 1994
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Electrode Structure. The electrodes were prepared by sequentially depositing four layers: the sensing layer (l), the insulating layer (2), the interference-eliminating layer (3), and the biocompatible layer (4). Figure 1B is a schematic drawing of the sensor. The thin polymer layers were well protected by containment within the polyimide sleeve. The sensing layer was made by “wiring” rGOx to the gold electrode through a redox hydrogel to which the enzyme was covalently bound. The electrodes were prepared as follows: 10 mg/mL solutions were made from (1) the PVI-Os redox polymer in water, (2) the cross-linker, PEGDGE, in water, and (3) the enzyme, rGOx, in a 10 mM HEPES solution adjusted topH8.15. Theredox hydrogelwas formed bymixing the three solutions so that the final composition (by weight) was 52% redox polymer, 35% enzyme, and 13% cross-linker. The insulating layer prevented electrical contact between the redox hydrogel and the interference-eliminating enzymes (HRP and LOX). PAL-PAZ was used as the insulating material. The film was deposited from a solution obtained by mixinginvolumeratioof 1:1, 1:2,or 1:3,aPALsolution (4.5 mg in 100 mM HEPES buffer at pH 7.0) and a freshly prepared PAZ solution (30 mg/mL). The PAZ solution was used within 15 min of preparation. The interference-eliminating layer was prepared according to a previously published protocol.3s A 50-pl aliquot of a 12 mg/mL freshly prepared sodium periodate solution was added to 100 pl of a solution containing 20 mg/mL H R P (HRP-VI or HRP-BM) and 100 mg/mL LOX (LOX or rLOx) in 0.1 M sodium bicarbonate, and the mixture was incubated in the dark for 2 h. Alternatively, the oxidation of H R P was carried out prior to adding LOX and cross-linking. The biocompatible layer films were photo-cross-linked by exposure to UV light (UVP, Inc., San Gabriel, CA; BlakRay; spectral peak at 360 nm, UV irradiance at the sample 200 mW/cm2) for 1 min. The initiator used was 2,2dimethoxy-2-phenylacetophenone (Aldrich). A solution of 300 mg/mL of the initiator in 1-vinyl-2-pyrrolidinone (Aldrich) was added to the prepolymer mixtures. Approximately 30 pL of the initiator solution was added per milliliter of 10% (w/w) aqueous solution of the tetraacrylated PEO. The prepolymers were cross-linked in situ inside the recess of the electrode. The films were prepared by filling the recess with the prepolymer solution twice and exposing the electrode to the UV light source after each time the cavity was filled. In vitro experiments were carried out in batch fashion at 25 and 37 OC, using a conventional three-electrode electrochemical cell with the enzyme-modified gold wire as the working electrode, a platinum wire as the counter electrode and a saturated calomel reference electrode (SCE). The electrolyte was a 20 mM phosphate-buffered saline solution containing 0.15 M NaCl at pH 7.15. Experiments in serum were performed at 37 O C , adding 100 pL antibioticantimycotic solution to 10 mL of serum. Phosphate-buffered saline and serum were agitated during the experiments. The working potential was +0.3 V vs SCE for experiments with the PVI-Os polymers. In vivo experiments (6-10 h) were carried out in 300-g male SpragueDawley rats. The rats were fasted overnight and prior to the experiment were anaesthetized with an intraperitoneal (i.p.) injection of sodium pentobarbital (65 3134
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mg/kg rat wt). An i.p. injection of atropine sulfate (166 mg/ kg rat wt) was then administered to suppress respiratory depression. Once the rat was anaesthetized, a portion of the rat’s abdomen was shaved and coated with a conductive gel, and an Ag/AgCl surface skin reference electrode was attached. Sensors were then implanted subcutaneously, using a 22-gauge Per-Q-Cath Introducer (Gesco International, San Antonio, TX), on the rat’s thorax or subcutaneously in the intrascapular area through a small surgical incision. The sensors were taped to the skin to avoid sensor movement. The sensors, along with the reference electrode, were connected to an in-housebuilt bipotentiostat. The operating potential of the sensors was 0.3 V vs Ag/AgCl, with the Ag/AgCl electrode serving as both the reference and counter electrodes. Sensor readings were collected using a data logger (Rustrak Ranger, East Greenwich, RI) and at the end of the experiment were transferred to a computer. During the experiment, the rat’s body temperature was maintained at 37 OC by a homeostatic blanket. The sensors were allowed to reach a basal signal level for at least 1 h before blood sampling was started. Blood samples were obtained from the tail vein, and all blood samples were analyzed by use of a glucose analyzer (YSI, Inc., Yellow Springs, OH; Model 23A). Approximately 30 min after the start of blood sampling, a i.p. glucose infusion was started using a syringe pump (Harvard Apparatus, South Natick, MA) at a rate of 120 mg of glucose min-I (kg of rat wt)-I. The glucose infusion was maintained for 1 h. At the end of the experiment, the rat was euthanized by sodium pentobarbital injection i.p. or asphyxiation by C02, consistent with the recommendations of the American Veterinary Association. All animal experimentation was approved by the University of Texas Institutional Animal Use and Care Committee.
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RESULTS AND DISCUSSIONS Structure of the Sensor. The Recessed Electrode. The recess, i.e., channel, in the polyimide-insulated wire was formed by electrochemical etching of the gold under galvanostatic control. By controlling the charge, the total amount of gold electrooxidized and dissolved as Au(CN)2- was defined. When the conditions were set so that the CN- transport into the channel and the Au(CN)2- transport out of it were not rate limiting, a flat gold wire surface was produced at the bottom of channels with aspect ratios of 0.5-2.0. Thus, when the CN- concentration was high enough and the wires were ultrasonically vibrated, the tips of gold wires were flat. Passage of 1.5 C/electrode at 8 mA current/electrode produced 125pm-deep cavities. At theoretical efficiency for one-electron oxidation, 3.08 mg of gold would have been etched. The amount of gold actually etched was only 0.076 mg, showing significant CN- or water oxidation. Nevertheless, the process was reproducible, accurate, and fast, with 20 electrodes being processed in each batch in less than 5 min. Structure and Performance. The depth of the channel and the thickness of the polymer layers in it controlled the mass transport of glucose to the sensing layer. By controlling these parameters, the apparent K,,,was adjusted to -20 mM glucose. The polyimide wall of the channel also protected the four polymer and polymer-znzyme layers against mechanical damage and reduced the hazard of their loss in the body.
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Glucose Conc. (mM)
Figure 2. Dependence of the sensitivity on the Os loading of the wire: (V)PVI,-Os and (V)PV15-Os.
Because the glucose electrooxidation current was limited by glucose mass transport through the recess and its polymer films, rather than by mass transport to the tissue-exposed tip, the current was practically insensitive to motion. Evidently, the electrooxidation rate of glucose in the recessed sensing layer was slower than the rate of glucose diffusion to the channel’s outer fluid contacting interface. PVI5-Os was chosen as the wire of the sensing layer. The subscript 5 is used to indicate that, on the average, every fifth vinylimidazole mer carried an electron-relaying osmium center.50 The results for PVI5-Os and for PVI3-Os (every third vinylimidazole mer carrying an osmium center) are compared in Figure 2, showing higher current density of glucose electrooxidation on electrodes made with PVIs-Os than on those made with PVI3-Os. Depth of the Recess and the Sensing Layer. Channels of 125- and 250-pm depth were investigated to assess the dependence of the current on the depth of the recess (Figure 3), with the total amount of PVIs-Os and rGOx being kept constant. Much of the loss in current in the deeper cavities resulted not from reduced glucose mass transport but from adsorptive retention of part of the enzyme and polymer on the polyimide wall when microdrops of the component solutions were introduced into the recess in the process of making the electrodes. Through repeated rinsing with water, some of the adsorbed polymer and enzyme on the walls were washed onto the electrode surface, increasing the current. The highest currents were seen after five washings. When the thickness of the sensing layer was increased through increasing the number of coatings (Figure 4), the current reached a maximum and then dropped. For the preferred 125-pm recess, 10 coatings, producing a 13-pm-thick wired-rGOx sensing layer, yielded sensors that had the desired characteristics for in vivo use. The Insulating Layer. This layer electrically insulated the redox enzymes of the interference-eliminating layer (HRP and LOX) from the wired rGOx layer. PAL cross-linking with PAZ, forming a polycationic network at pH 7.0, was
-
1
1
1
1
1
1
1
1
1
1
1 0 15 2 0 2 5 3 0 3 5 40 45 50 5 5 60 6 5 70 75 80 8 5
Glucose Conc. (mM)
Figure 3. Dependenceof the sensitivity on the recess depth: (0)125and (0)250-pmdeep electrodes modified with the same amount of sensing layer; E = +0.3 V vs SCE; 20 mM phosphate buffer containing 0.15 M NaCi at pH 7.15 and 25 OC. 140
I
I
I
I
120
100
0
80
: v
0 >
60
40
20
0
1
I
2
4
I
6
6
10
1
1
I
12
14
16
18
20
Sensing Layer Thickness (p”
Figure 4. Dependence of the normalized current on the thickness of the sensing layer: 20 mM phosphate buffer containing 0.15 M NaCi at pH 7.15 and 37 OC; E = +0.3 V vs SCE. Q Is the number of coulombs of Os redox centers in the layer.
used. The best results were obtained using 1:2 PAL-PAZ (Figure 5 ) , with three coatings applied to form a -7-pmthick cross-linked film. The Interference- Eliminating Layer. The interferants, particularly ascorbate, urate, and acetaminophenol, were oxidized in the third layer, containing LOXand HRP. In this layer, lactate, the typical concentration of which in blood is 1 mM, reacted with 0 2 to form HzOz and pyruvate. H202, in the presence of HRP, oxidizes ascorbate, urate, and acetaminophenol, being reduced to ~ a t e r . The ~ ~ preferred , ~ ~ coimmobilization process involved two separate steps: periodate oxidation of oligosaccharide functions of HRP to aldehydes, followed by mixing with LOX and formation of multiple Schiff bases between HRP aldehydes and LOXamines (e.g., lysines) and between H R P aldehydes and amines. The AnalyticalChemistty, Vol. 66, No. 19, October 1, 1994
3135
I
Table 1. Sensoor Characterlrtks'
l o9 o0
60
e
j(pA/cm*)
EH
LB
tr(s)
current variance(%)
33.9
69.1
18.5
33.4
30-90
5.0
i is the current measured at 37 OC and at 10 mM glucoseconcentration; j i s thecurrentdensitymeasuredat 37 'Cat 10mMglucoseconcentration; Kmapp is the apparent Michaelis-Menten coefficient determined from an electrochemical Eadie-Hoffstee (EH) or a Lineweaver-Burk (LB) plot; 1, is the 10-902 rise time, 90 s for 2 mM and 30 s for 20 mM glucose concentration; current variance is the maximum deviation from the mean value measured during the 72-h test conducted in 10 mM glucose in the presence of interferants. The current was monitored continuously at 37 OC.
v
Z
KmpPP(mM)
i(nA)
50
40
i 01
I
0
10
20
40
30
50
,
1
60
70
80
I
100
90
Time ( h )
Figure 5. Stability of the sensitivity of electrodes varying in sensing layer thickness and Insulating layer composition and thickness. (0) 5-pm PV15-Os, 4-pm PAL PAZ (ratio 1:l): (0)7.5-pm PVIs-Os, 4-pm PAL PA2 (ratio 1:l): (V)7.5-pm PV15-Os, 4.5-pm PAL PA2 (ratio 1:2):and (V)12.5-pm PV15-Os, 7-pm PAL PA2 (ratio 1:2): 20 mM phosphate buffer containing 0.15 M NaCl at pH 7.15, 37 OC, 10 mM glucose; +0.3 V vs SCE.
+
+
+
+
Stability and Other Characteristics. In order to improve the stability, the more thermostable rGOx rather than GOx49 was used and glucose transport was reduced to make the sensor current diffusion, not enzyme turnover, limited. The glucose flux was attenuated by the three outer layers and the sensing layer itself. Because the sensing layer had a large excess of glucose oxidase, its activity greatly exceeding that needed for electrooxidizing the attenuated glucose flux, the sensor's stability was improved. The stabilitywas tested in the presence of 0.1 mM ascorbate in 10 mM glucose at 37 OC. The current output of a typical optimized electrode was 35 nA and the apparent K m , derived from an Eadie-Hofstee plot, was 20 mM (Table 1). The 10-90% response time was 1 min. As expected, and as can be seen in Figure 5, with thinner films the glucose mass transport increased, Le., the current was higher, while for thicker films the stability improved. Because of the high sensitivity of thin sensing film (- 1 pm) electrodes (>10-*A cm-* M-l), a 1 order of magnitude decrease in sensitivity could be traded for stability, while the currents remained high enough to be easily measured. As seen in Figure 5, the sensitivity of the stabilized sensors did not change by more than f5% for 72 h of operation at 37 OC. After a small initial decrease in sensitivity, it increased to a maximum after 40 h, and the final 72-h sensitivity was almost identical with the initial. There was no measurable difference in the apparent stability of the sensitivity of continuously operated and resting electrodes measured at 6-h intervals. The characteristics of the electrodes are summarized in Table 1. Each entry represents an average value for five electrodes. The baseline currents were typically
