Development of a novel enzyme based glucose sensor

S.B. Petersen, B. Svensson, and S. Pedersen (Eds), CarbohydrateBioengineering 9 Elsevier Science B.V. All rights reserved.

49

Development of a novel enzyme based glucose sensor F. Spener, R. Steinkuhl, C. Dumschat, H. Hinkers, K. Cammann and M. Knoll Institut ftir Chemo- und Biosensorik Mtinster, Mendelstr.7, D-48149 Mtinster, Germany

Abstract

Glucose oxidase from Aspergillus niger is an extremely stable, FAD-dependent enzyme that has found wide application in electrochemical sensing of glucose in blood. Conventional sensor technology immobilises this enzyme in membranes deposited on top of the transducer. Here we report a new containment technology where the glucose oxidase is immobilised in gelatine or polyvinylalcohole and is deposited in the chip in pyramidal containments produced on silicon by anisotropic etching. This configuration of the sensor enhances the adhesion and stability of the membrane and has been applied to monitoring glucose in serum discontinuously as well as continuously in a flow-through system. Moreover we demonstrate the applicability of the containment technology to monitor glucose ex vivo with the help of a microdialysis system.

1. INTRODUCTION 1.1. Glucose oxidase

Glucose oxidase (I]-D-glucose: oxygen 1-oxidoreductase, EC 1.1.3.4) is a FAD-dependent enzyme that catalyzes the oxidation of ~-D-glucose by dioxygen to hydrogen peroxide and 8gluconolactone, which subsequently hydrolyzes spontaneously to gluconic acid. Apart from glucose monitoring in food, drinks and fermentation processes the most important application of this enzyme is in clinical diagnostics. Here it is used as a component of colorimetric diagnostic kits, of dry reagent test strips, and more recently of biosensors for the determination of glucose in blood, serum of plasma. Glucose oxidase (GOD) exhibits useful properties with respect to widespread application in biosensors because of the high specificity for glucose and an extremely high stability [1]. The enzyme has been isolated from various moulds, from red algae, citrus fruits, insects and bacteria. The most widly used enzyme in terms of research and commercialized products, however, is that from Aspergillus niger, a glycoprotein with a high-mannose type carbohydrate content of 10 to 16 % of its molecular mass. The carbohydrate component appears to be in form of a branched polysaccharide that partially surrounds the protein core [2]. Many of the enzyme's physical properties, such as high solubility in water and resistance to proteases for example, may be ascribed to its carbohydrate shell. Partial removal reduces the enzyme's stability.

50

Figure 1. Tertiary structure of FAD-containing glucose oxidase (monomer) from

Aspergillus niger [3]. The holo-enzyme is a homodimer, where each subunit contains the tightly bound (Kd = 1 x 101~ but not covalently attached coenzyme (FAD), and the reaction mechanism is essentially

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Figure 2. Natural and artifical substrates in redox reactions catalyzed by glucose oxidase from

Aspergillus niger.

51 dependent on this coenzyme. The recently unraveled tertiary structure of glucose oxidase (Fig. 1) from A. niger [3] shows that this enzyme allows electron acceptors other than dioxygen to reoxidize the reduced FAD in the same way. This is important for the application of the enzyme in the amperometric approach to glucose sensors, because the transfer of electrons between the active site of the reduced enzyme and an electrode takes place slowly or not at all. Transfer of electrons can be facilitated if a small electron acceptor is used as a mediator [ 1]. These compounds have the advantage that they allow amperometric biosensors to operate at relatively low potentials, and this can lead to a decrease in interference from electroactive compounds (Fig. 2).

1.2. Biosensors Given the extraordinary operational stability of glucose oxidase and the world-wide high demand for monitoring glucose in the blood of patients with type I diabetes (insulin-dependent diabetes mellitus) as a parameter for insulin injections that compensate the defect in insulin production by the pancreas, it is no surprise that the foremost application of biosensors to date is in the medical field. Initially biosensors for monitoring blood glucose have been used as analyzers in clinical laboratories. The US-company Yellow-Springs Instruments was the first to offer commercial glucose analyzers, based on the patent of L.C. Clark. Today the decentralized employment of point of care diagnostics in the doctor's office or in intensive care units and the bed-side measurements in hospitals are of greater interest (for example the multiparameter sensor of ISTAT Corp. Princeton, USA). Of utmost importance today, however, is self-testing by patients using miniaturized electrodes with immobilized GOD for blood glucose sensing. In this area 3 sensor systems are in different states of development. A big success already are the commercially available, disposable pen- or card-type sensors. The more provocative approach aims at continuous monitoring of blood glucose to attain normoglycemia and to avoid acute metabolic disturbances. With this in mind work on the development of implantable biosensors is going on for years. Due to the - in this case still limited - stability of the enzyme and still not optimal biocompatibility of all the sensor materials used, this concept has not yet been realized. A solution in the far future may be artificial organs, such as the pancreas, with an inbuild glucose sensor and insulin pump in nanotechnology. The more realistic concept for realization in the near future is a continuously working microdialysis system, were dialysates of the subcutaneuos tissue are channeled to a sensor for continuous glucose monitoring ex vivo [4]. Our approach to the glucose sensor is based on the GOD-catalyzed oxidation of glucose to gluconolactone and concomitant reduction of dioxigen to H202. The latter is electrochemically oxidized at a platinum anode. The anode is polarized at + 600 mV vs. an Ag/AgC1 reference electrode. Often the enzyme is immobilized in a membrane coveting the surface of the anode. The current depends linearly on the glucose concentration in the solution [5]. Until now most enzymatic biosensors have a membrane fixed on top of the transducer. This method often leads to malfunction of the sensor arising from problems like inadequate membrane adhesion and insufficient mechanical stability. Dipcoating procedures are difficult to perform in a reproducible way in order to obtain sensors with identical performances. Sensors with membranes that have separately been casted and mounted to an electrode can hardly be miniaturized and are therefore not suited for potential implantation.

52 In order to solve these problems we developed a new concept for membrane deposition [6, 7], the so called c o n t a i n m e n t s e n s o r s on the basis of micromechanically etched cavities in silicon substrates (<100> crystal orientation, 380 lam thick) with openings towards the analyte solution between 120 and 480 lam width, where the membrane is not located on but in the chip (Fig. 3). At the sides of the cavities platinum electrodes are deposited by means of semiconductor technology such as photolithography and physical vapour deposition. Here we demonstrate the application of this technology to the sensing of glucose in body fluids.

Pt Si SiO 2

Containment with immobilized enzyme

Figure 3. Schematic view of the containment sensor.

2. EXPERIMENTAL

2.1. Chip fabrication A silicon wafer (p-type, 3-inch, (100)-surface, 380 l.tm thick, optically polished on both sides) was oxidized in wet ambient at 1200 ~ for 210 min to get a 1.5 l.tm thick oxide layer. Spin-coating was performed sequentially with HMDS (hexamethyldisilazane) as adhesion promoter and photoresist (AZ 5214E/Hoechst) at 4000 rpm followed by a prebake at 90 ~ for 5 min. The resist layer had a thickness of 1.4 pm. It was exposed to about 80 mJ/cmz of mid UV radiation through a photomask to define the containments on the front side of the wafer. Development with AZ 524MIF/Hoechst was followed by a postbake at 120 ~ for 60 s. A protective coating was then applied to the back side of the wafer to withstand the following oxide etching process which was carried out with BOE (buffered oxide etchant) to get quadratic holes in the front side. Photoresist and protective layers where removed and silicon was etched anisotropically in 20 % KOH solution at 75 ~ for 6 h to create the containment holes. The silicon oxide was then stripped with hydrofluoric acid and a new 150 nm thick oxide layer was created thermally in dry oxygen at 1050 ~ In order to form platinum electrodes the front side of the wafer was first sputter deposited with a 1/am thick aluminum layer which acts as a sacrificial layer for reliable platinum patterning in a lift-off process. Using our newly developed electrospray-coating technique the aluminum layer was subsequently covered with a photoresist layer, again AZ 5214E from

53 Hoechst, which was now used in image 2 reversal mode. The prebake at 90 ~ for 5 min was sequentially followed by a 40 mJ/cm UV exposure through a photomask, a reversal bake at 120 ~ for 5 min, a 320 mJ/cm UV flood exposure, and development with AZ 524MIF. Then aluminum etching was performed with phosphoric acid etchant (PES 83.5-5.5-5.5/Merck) to get a negative mask for the following application of the platinum electrodes and conducting lines. A 200 nm thick platinum layer was then sputter deposited and patterned by removing the aluminum layer in a sodium hydroxide solution. Anodic bonding of a Pyrex | was performed at 500 ~ and 300 V for 60 min. Finally the chips were separated.

glass wafer

2.2. Enzyme immobilisation and membrane deposition The buffer (pH 7.0) was prepared by solving 8.1 g NaC1 (Merck), 0.272 g KH2PO 4 (Sigma), and 0.697 g K2HPO 4 (Sigma) in 1 1 deionized water. GOD (200 U/mg) was obtained from Fluka. 50 mg of acid photo gelatine (Filmfabrik Wolfen, Germany) were allowed to swell in 0.5 ml deionized water for 60 min at room temperature. Then 0.5 ml of the prepared buffer (pH 7.0) solution were added. After mixing the solution was kept at 35 ~ for 60 min, then different amounts of GOD-powder were added to 100 lal of the warm stirred gelatine solution in a small Eppendorf vessel. The vessel with the enzyme was placed in a chamber and the transducer chip was dipped into the enzyme-solution to cover the containment opening with it. Then the chamber was evacuated in order to fill the containment with the gelatine solution. After gelling the sensor surface was cleaned and the sensor was examined under a microscope to ensure a properly filled containment. Then the sensor was stored in buffer solution until the first measurement. Furthermore polyvinylalcohol functionalized with stilbene groups (PVA-SbQ) was used for enzyme immobilization [8]. 1600 U GOD were dissolved in 5 0 0 m g PVA-SbQ 500-88 solution. The solution was filled into the containment and then the PVA-SbQ was hardened by UV-irradiation. 2.3. Measurements The measurements were performed in a two electrode configuration. A chloridized silver wire was used as reference electrode. A voltage of 0.6 V was applied between the containment electrode and the Ag/AgC1 reference electrode (potentiostat Autolab PSTA 1 0 / 4 channel, Eco Chem). The sensors were tested in the buffer solution described above. Before starting calibration the sensors were polarized until the current was lower than 1 hA, which took about one hour. Known amounts of a highly concentrated glucose solution (Merck) were added to the stirred buffer solution to obtain the calibration graphs. When not in use the sensors were stored in buffer solution at 4 ~ The sensor was also probed in a flow system (fixed in a block of acrylic glass with a dead volume of approximatly 10 laD, equipped with a pump from Meredos and a valve from Knauer were used (the flow rate was 50 pl/min). The response time was obtained from the response curves in the flow system. The stability of the sensor was determined by immersing the polarised sensor together with the chloridized silver wire in stirred undiluted human serum (purchased from Sigma) and recording the current. The concentration changes in the undiluted serum were obtained by adding known amounts of high concentrated glucose solution.

54 3. RESULTS AND DISCUSSION The response curves and the corresponding calibration graphs of simultaneously measured containment sensors with different opening sizes (120, 250, 480 pm length of a side) toward the analyte solution are shown in Fig. 4. All sensors with gelatine based membranes responded fast and stable on changes in the glucose concentration. As expected the sensitivity of the sensors increased with increasing opening size. The detection limit was lower than 0.05 mM.

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In Fig. 5 it is demonstrated that the sensor is well suited for a flow system configuration. The sensor responded linearly and reproducibly in a concentration range relevant for continuous subcutaneous monitoring of glucose in a microdialysis system, if the microdialysis needle and flow rate are chosen properly. The response time (t90) in the flow system (Fig. 5) is 65 s for the change in concentration from 0.5 to 1.25 mM, and 104 s for the change from 1.25 to 0.5 mM.

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56 As demonstrated in Fig. 6, the sensor shows a stable signal in undiluted human serum over more than two days. This demonstrates the applicability of the sensor in undiluted body fluids. Because the linear range of the gelatine based membranes was not sufficient in undiluted probes, the photosensitive PVA-SbQ was chosen for the immobilization of GOD. The calibration plot indeed shows a wider linear range (Fig. 7). 100

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It is possible to store the sensors with GOD immobilized in PVA for more than 3 months in the refrigerator with a minute decline in sensor characteristics only. The influence of interfering substances was not investigated but a behavior similar to planar sensors using the same materials is expected. The results show that enzyme membrane deposition in silicon containments makes glucose sensors with good analytical response behaviour feasible. Linear range, in particular the normal physiological range of 3.9- 6.1 mM (70- 100 mg/dl) for glucose in blood and above in the diseased state, as well as lifetime and response times are reasonable well and acceptable for most practical applications, also for use in flow systems combined with microdialysis sampling. In the introduction we alluded to the lack of stable, implantable and long lasting glucose sensors [9, 10]. On the way toward this goal one immediate solution with application potential is the development of a miniaturized microdialysis system. Such a microdialysis system under development in our laboratory is shown in Fig. 8.

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The analytical response behaviour of these containment sensors is comparable with other sensor configurations which are fabricated in silicon planar technology, but offers the following advantages: - Containment sensors are very robust because the mechanically sensitive membrane is protected inside the chip against damage. Also adhesion problems with the membranes are minimized since they are attached inside the containment. Membrane patterning is not needed. - Encapsulation is no longer a problem since the main components of the sensor (membrane and metal electrodes) are located inside the containment. Encapsulation is further simplified -

58 due to the fact that the conducting leads and the sensitive sensor surfaces are located on opposite sides of the wafer. Thus, encapsulation can be performed on the top side of the whole wafer. - Containment sensors are ideally suited for integration into microsystems as the dead volume added to the system is minute. - The fabrication process is entirely compatible with mass-production technologies since it is a full wafer process. In mass production even the membrane deposition can be performed as full wafer process either under vacuum or with an automated dispense system. Only soldering the plug must be done after the chips have been separated. In conclusion, the containment concept allows the fabrication of glucose sensors with good analytical response behaviour and with a considerably improved technology.

4. REFERENCES

1 2 3

R. Wilson and A. P. F. Turner, Biosens. & Bioelectron., 7 (1992) 165. J.H. Pazur, K. Kleppe and A. Cepure, Arch. Biochem. Biophys., 111 (1965) 351. H.J. Hecht, H. M. Kalisz, J. Hendle, R. D. Schmid and D. Schomburg, J. Mol. Biol., 229 (1993) 153. 4 C. Meyerhoff, F. J. Mennel, F. Bischof, F. Sternberg and E. E. Pfeiffer, Horm. Metab. Res., 26 (1994) 538. 5 F. Scheller and F. Schubert, B iosensoren. Akademie-Verlag, Berlin, 1989. 6 M. Knoll, German Patent DE 4115414A1 (1991). 7 R. Steinkuhl, C. Sundermeier, C. Dumschat, K. Cammann and M. Knoll, German Patent application Nr. P 4337418.2 (1993). 8 R. Renneberg, K. Sonomoto, S. Katoh and A. Tanaka, Appl. Microbiol. Biotechnol., 28 (1988) 1. 9 G . S . Wilson, Y. Zhang, G. Reach, D. Moatti-Sirat, V. Poitout, D. R. Th6venot, F. Lemonnier and J.-C. Klein, Clin. Chem., 38 (1992) 1613. 10 S. J. Updike, M. C. Shults, R. K. Rhodes, B. J. Gilligan, J. O. Luebow and D. von Heimburg, ASAIO J., 40 (1994) 157.