Glucose biosensor based on a reagentless graphite-epoxy screen-printable biocomposite

Sensors and Actuators B 45 (1997) 55 – 62

Glucose biosensor based on a reagentless graphite-epoxy screen-printable biocomposite Carlos A. Gala´n-Vidal a,c, Javier Mun˜oz b, Carlos Domı´nguez b, Salvador Alegret a,* a

Departament de Quı´mica, Uni6ersitat Auto`noma de Barcelona, Grup de Sensors i Biosensors, E-08193 Bellaterra, Spain b Departamento de Microsistemas y Tecnologı´a de Silicio, Instituto de Microelectro´nica de Barcelona, Centro Nacional de Microelectro´nica. C.S.I.C, E-008193 Bellaterra, Spain c Seccio´n de Electroquı´mica, Uni6ersidad Auto´noma de Hidalgo, Hidalgo, Mexico Received 12 August 1996; received in revised form 4 August 1997; accepted 8 August 1997

Abstract New screen-printable graphite-epoxy composites are developed. Different formulations of this electrochemical sensing material were prepared in order to construct amperometric transducers by thick-film technology. Resulting devices were characterized electrochemically using cyclic voltammetry and an optimal graphite content was also determined. The applicability of the optimized transducer material has been evaluated through electrochemical detection of hydrogen peroxide and by the development of a glucose biosensor. In this case, the screen-printable composite is bulk-modified with the addition of glucose oxidase (GOD). The detection of hydrogen peroxide at 1150 mV (Ag/AgCl) serves as the analytical signal. The sensor shows a linear response range for glucose between 0.05 and 2.5 mM in a pH 7.0 buffered solution with 0.1 M phosphate and 0.1 M KCl. Other biosensing systems can be transferred to thick-film technology based on the use of analogous materials (biocomposites) previously studied in our group, i.e. acetylcholinesterase, peroxidase, alcohol oxidase, etc. © 1997 Elsevier Science S.A. Keywords: Amperometry; Biosensor; Glucose; Screen-printing; Biocomposite; GOD

1. Introduction A growing interest in the development of thick film technology applied to chemical sensors in order to realize small-size transducers has been increasing as an alternative to standard electrodes for some special applications, such as environmental, food industry, and biomedicine. These transducers are being developed to obtain chemical sensors and biosensors that can be used in a short time and in small volumes, being fabricated on a large scale and at low cost [1 – 3]. In the case of biosensors that include enzymes or immunological species, the immobilization procedure is the most critical aspect of the whole fabrication process. These methods have to be performed by suitable techniques depending on the specific performance. Amperometric biosensors developed in planar structures have to combine enhancement of the selectivity, sensitivity and application * Corresponding author. Tel.: +343 581 1017; fax: +343 581 2477. 0925-4005/97/$17.00 © 1997 Elsevier Science S.A. All rights reserved. PII S 0 9 2 5 - 4 0 0 5 ( 9 7 ) 0 0 2 7 0 - 0

range together with the mass production capability of thick-film technology. On the other hand, polymer technology offers a wide range of materials that can be used in a variety of transducers. These possibilities open a new generation of rigid conducting composites that have been applied in the realization of biosensors. New composites based on incorporating the biocatalyst component into carbonbased and other types of conducting pastes (biocomposites) [4–10] are being formulated. Although most of the work is being devoted to immobilize the enzymaticbased recognition elements after the fabrication of the transducer which implies the application of several layers on the fabrication process [3]. There are technological drawbacks that make it difficult to compatibilize biocomposite formulations with the possibility to be screen printed on miniaturized transducers. Nevertheless a new generation of transducers based on a mixture of composite enzyme, and a curing process at lower temperatures to assure enzymatic activity are simplifying the development of solid-state chemical sensors.

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In the present work, suitable screen-printable graphite-epoxy pastes have been assayed as a transducer material. Subsequent modification of this material has resulted on a new type of biocomposite containing glucose oxidase (GOD), which has been applied to perform thick film glucose biosensors. The goal of this work is to develop a new glucose biosensor based on a reagentless biocomposite in order to simplify the immobilization step on the transducer.

Table 1 Screen printing paste recipes for transducers (T) and biosensors (B) Planar electrodes

T5 T10 T15 T20 B

ml/g(Paste)

Paste components Epoxy

Graphite

GOD Ciclohexanone

95 90 85 80 78.4

5 10 15 20 19.6

— — — — 2

40 80 150 200 250

2. Experimental

2.1. Reagents and solutions All reagents were commercially available and were employed without further purification. NaH2PO4 · H2O, Na2HPO4 · 2H2O, H2O2, K4[Fe(CN)6] · 3H2O, D(+ )Glucose · H2O were purchased from Merck and KCl from Fluka. All solutions were prepared with bidistilled water. Printing biocomposite were prepared using graphite powder (Aldrich) with a particle size of 1 – 2 mm, epoxy resin (Epotek H77, Epoxy Technology, Billerica, MA), GOD (type VII, EC 2.1.3.4 from Aspergillus niger) (Sigma) and cyclohexanone (Merck). The supporting electrolyte was an aqueous solution buffered at pH 7.0 with 0.1 M phosphate and 0.1 M KCl. Hydrogen peroxide solutions were prepared daily and standardized by a volumetric method [11]. Glucose solutions were prepared from benzoic acid saturated solution [12] and prior to use glucose was allowed to mutarotate for at least 1 day.

arrays of ten electrodes. The graphite-epoxy pastes were prepared by dispersing graphite powder in Epo-Tek H77 epoxy resin. Five different mixed proportions were prepared by varying the amount of graphite (Table 1). Once homogenized the composite’s viscosity was adjusted with cyclohexanone. Used a polyurethane squeegee, each of the resulting composite pastes were printed onto a glass fibre circuit board on which ten copper lines 1× 30 mm were patterned by a conventional photolithographic process. The printed layer was cured for 72 h at 40°C. Finally the copper lines were covered by a layer of epoxy diacrylate (Ebecryl 600, UCB Chemicals) and exposed through a mask under UV. Thus leaving a graphite-epoxy working area of 12 mm2 and a contact pad (Fig. 1). The transducers were stored dry at room temperature. The same transducer fabrication process explained above was used to glucose biosensor construction. The GOD-graphite-epoxy paste was prepared by simple dispersing graphite powder and GOD in Epo-Tek H77 epoxy resin. Once homogenized the biocomposite be-

2.2. Instrumentation Cyclic and linear voltammograms were obtained with a PGSTAT 20 Autolab potentiostat (Ecochemie). Three electrodes formed the electrochemical cell; a platinum counter electrode (Ingold), double junction Ag/AgCl reference electrode (Orion) with 0.1 M KCl external solution and either graphite-epoxy screen printing transducer or biosensor as working electrodes. Calibration curves were obtained with a 641 amperometric VA detector (Metrohm) coupled to a E506 Polarecord plotter (Metrohm). All measurements were performed at room temperature. The surface profile was tested with an Alpha-Step 200 profilemeter (Tencor).

2.3. Construction of the amperometric transducer and glucose biosensor. A semi-automatic screen printer (Marprint 350 Marbay, S.L., Barcelona, Spain) and a conventional monofilament polymide mask (200× 400 mm, 110 mesh, 70 mm thick) were used for constructed planar

Fig. 1. Schematic drawing and cross-sectional view of either the thick-film amperometric transducer or the amperometric GOD biosensor.

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fore printing, the viscosity was adjusted with cyclohexanone (Table 1). The printed biosensors were stored dry at 4°C. The resulting glucose biosensor based on the prepared biocomposites was characterized through the selective biocatalytic generation of hydrogen peroxide by the enzyme GOD. The H2O2 produced was measured amperometrically by direct oxidation on the surface of the biocomposite material electrode when a potential E was applied

3. Results and discussion

3.1. Electrochemical beha6iour of transducers The electrochemical characteristics of different types of amperometric transducers were evaluated from the cyclic voltammograms obtained for the Fe(CN)36 − / Fe(CN)46 − system were evaluated. Fig. 2a shows high potential position of the oxidation peak of Fe(II). The reduction peak in the studied potential range was unlocalized, suggesting a slow electron transfer. Nevertheless, it is observed that by increasing the content of graphite up to 20%, electrochemical behaviour is highly improved. Furthermore, this amount was the limit of the content of graphite since major problems related to composite printing and post-curing steps were found. In order to increase the kinetic of the electron transfer process, a manual mechanical polishing process by means of a 3 mm alumina paper (301044-001 polishing strips, Orion) on the transducer array was applied, removing a ca. 20 mm of composite layer (Fig. 3). The oxidation peak was shifted to lower potentials obtaining higher response currents from all transducers. This means that a faster electron transfer takes place, which by appearing the reduction peaks together with a lower separation between both anodic and catodic peaks was corroborated. The transducer containing 20% of graphite (T20) presented the best response and lowest overpotential. On the other hand, transducers containing less graphite such as 15 and 10% were respectively showing a decrease of 10 and 25% on sensitivity with respect to T20. However, a wider range of modification than T20 can be permitted for both T15 and T10 transducers, being more suitable formulations for certain applications (Fig. 2b) [13]. As expected from results obtained with polishable composites [14], this common treatment has been also implemented in the development of planar thick film transducers [15] involving a more complex technology. Nevertheless, this drawback has been solved by applying a surface activation process via electrochemical method compatible with the set-up of measurements of the amperometric transducers [16,17]. The transducer was activated on applying cyclic potential between −

Fig. 2. Cyclic voltammograms for Fe(CN)46 − 2.4 mM at unpolished (a) and polished (b) thick-film amperometric graphite-epoxy transducers with different graphite weight content. The supporting electrolyte is a 0.1 M phosphate and 0.1 M KCl solution at pH 7.0. Scan rate is 50 mV s − 1.

0.25 and + 1.60 V vs. Ag/AgCl, shifting the anodic peak potential and enhancing the sensitivity with respect to the untreated transducers up to analogous values to those obtained by the polishing method (Fig. 4). In this work, the optimized transducer was composed by 20% of graphite (w/w) and was activated via electrochemical treatment. The response parameters of this base transducer was evaluated by electrochemical detection of H2O2, and subsequently tested by adding GOD into the composite in order to fabricate a thick film glucose biosensor.

3.2. E6aluation of the amperometric transducer. The optimum oxidation potential of hydrogen peroxide was determinated by linear-sweep voltammetry (Fig. 5). The typical plateau obseved at 1150 mV corre-

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Fig. 3. Surface profiles of (a) untreated T20 transducer and (b) polished T20 transducer. The stylus force is 10 mg. T20 stands for 20% (w/w) graphite.

sponds to the diffusional limiting current of H2O2 oxidation. This potential agrees with previous studies using similar polymeric matrix [14], being selected as working potential for the next experiments. As already known, rigid graphite composites present high oxidation potentials and a wide anodic potential window [18].

Fig. 4. Cyclic voltammograms for Fe(CN)46 − 2.4 mM at T20 transducer, for conditions see Fig. 2. T20 stands for 20% (w/w) graphite. Scan rate: 50 mV s − 1.

Fig. 5. Linear voltammograms for hydrogen peroxide at T20 transducer. The supporting electrolyte is a 0.1 M phosphate and 0.1 M KCl solution at pH 7.0. Scan rate is 5 mV s − 1. T20 stands for 20% (w/w) graphite.

The pH effect on the response of the transducer was determined by hydrodynamic amperometry in support electrolyte and 0.22 mM H2O2 (Fig. 6). A minimum influence on the residual current between pH 2 and 8 was observed, although an important increase from pH 9 was apparent. This behaviour takes place due to a shifting of the electrochemical equilibrium of the oxida-

Fig. 6. Effect of the pH on the signal of a T20 transducer in 0.1 M phosphate and 0.1 M KCl buffer solution at pH 7.0. Applied potential: 1150 mV vs. Ag/AgCl. T20 stands for 20% (w/w) graphite.

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The construction reproducibility was evaluated from relative standard deviation of the slope of H2O2 calibration curves. Between different transducers of the same batch it was 3% (n=10). The shelf life was evaluated from the slopes of different transducers of the same batch through 0.5 year and no significant change was shown; so the shelf life is greater than 6 months.

3.3. E6aluation of the amperometric glucose biosensor

Fig. 7. Calibration graphs for two T20 transducers of the same batch. Working solution: 0.1 M phosphate and 0.1 M KCl buffer solution at pH 7.0. Applied potential: 1150 mV vs. Ag/AgCl. T20 stands for 20% (w/w) graphite.

tion of water, as well as to an etching process on the composite surface [19]. This etching (verified using a magnifying glass) provides an increase on the porosity of the material rendering a transducer with higher effective area. The addition of H2O2 gave higher currents at increasing pH values when compared with currents obtained with the buffer electrolyte at the same pH values (Fig. 6). As a consequence of the results, pH 7.0 was fixed for further works since at this pH any kind of etching mechanism of the surface composite takes place, and higher sensitivities are found. The lower and upper limit of linear response (LLLR and ULLR, respectively) were found graphically from calibration graphs (Fig. 7). The linear range of the system was between 2×10 − 6 and 5 ×10 − 4 M of hydrogen peroxide with a slope of 31.990.7 mA mM − 1 (r 2 =0.9990). The ULLR was limited by the capability of the transducer to oxidize the analyte as it comes in contact with its surface. If 3 S:N ratio is considered, the limit of detection is 5.10 − 7 M. The drift of the residual current generated by the system is 4 nA min − 1. On the other hand, the 95% response time of the transducer was 12 s when the H2O2 concentration was increased from 0 to 2.2 × · 10 − 4 M. During ensuing calibrations, the signal was very stable between consecutive additions. For consecutive calibration curves of hydrogen peroxide between 2.2× 10 − 5 M and 5 ×10 − 4 M a slope relative standard deviation of 3% was found for a 95% confidence level (n= 5).

When GOD and glucose interact, hydrogen peroxide is produced at the biosensor surface. In this case, hydrogen peroxide can be amperometrically measured by direct oxidation on the surface of the biocomposite. The optimum oxidation potential was checked by linear-sweep voltammetry, being analogous to that employed with the unmodified composite. The final glucose electrodes were characterized for the detection of glucose by studying the pH influence on the amperometric response (Fig. 8). The analytical signal was increasing up to pH 7.5 and subsequently decreasing due to the loss of activity of the enzyme. The higher currents were explained as a combination between major enzymatic activity and sensitivity of transducers at increasing pH values. As a consequence, the optimal working pH for the biosensor was found to be pH 7.0, in which the composite was more stable. The LLLR and ULLR were found graphically from calibration graphs (Fig. 9b). The linear range of the system was ranged between 0.05 and 2.5 mM of glucose with a slope of 205 9 11 nA mM − 1 (r 2 \ 0.999), which is six times greater than reported for conventional

Fig. 8. Effect of the pH on the signal of a glucose biosensor. Working solution: 0.1 M phosphate and 0.1 M KCl buffer solution at pH 7.0. Applied potential: 1150 mV vs. Ag/AgCl.

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i=

imax × Cglucose K app m + Cglucose

(1)

where i is the current measured, Cglucose, glucose concentration and imax the maximum current. It was found a K app m = 7.0 mM being similar to the value obtained for a biosensor with the same biocomposite material, but without cyclohexanone [20]. The construction reproducibility was evaluated from the relative standard deviation of the slope of glucose calibration curves, resulting 8.2% (n= 10) between different transducers of the same batch. On the other hand, the shelf life was evaluated from the slopes of different biosensors of a same batch through 2 months and no significant change was shown. Good results have been obtained by applying high oxidation potential. However, interferences from other electroactive species at these voltages were found. Previous work has shown that a decrease in oxidation potentials can be achieved by adding into the composite formulation Au–Pd (900 mV) [21], tetrathiafulvalene (150 mv) [22], peroxidase and peroxidase-Pt (− 300 and −50 mV, respectively) [23]. In our research, new developments based on modification of biocomposites are being carried out, and new screen-printed amperometric biosensors are being tested.

4. Conclusions

Fig. 9. (a) Dynamic response of the glucose sensor for a concentration range from 0 to 2.4 mM; (b) corresponding calibration graph. Working solution: 0.1 M phosphate and 0.1 M KCl buffer solution at pH 7.0. Applied potential: 1150 mV vs. Ag/AgCl.

glucose biosensors with analogous biocomposite materials and detection principles [20]. The ULLR was limited by the substrate saturation of the enzyme related with Km. The drift of the residual current generated by the system is 5 nA min − 1 and is similar to the transducer. In this case, the 95% response time of the biosensor is 25 s when the glucose concentration increases from 0 to 1 mM. During ensuing calibrations the signal is stable between consecutive additions (Fig. 9a). The Michaelis–Menten apparent constant (K app m ) was determinated from saturation curves of calibration (between 0 and 100 mM of glucose) by non-linear regression uses the Marquardt – Levenberg algorithm. The equation used at the iterative process was:

New composites based on screen printable graphiteepoxy formulations have been developed. This material has been optimized and evaluated in order to construct stable amperometric transducers. The optimal graphite contents in the composite was 20% since a higher amount of graphite lead to difficult printing and curing processes. On the other hand, transducers that contained less graphite showed lower sensitivities. The optimized transducers can be activated after the construction process by applying two different methods in order to improve their electrochemical performance. The polishing method was more complex than the electrochemical one. Therefore, a surface activation step via the electrochemical method compatible with the setup of measurements of the amperometric transducer is preferred. The activation procedure was carried out by applying cyclic potential between −0.25 and +1.60 V vs. Ag/AgCl. The polymeric matrix has shown its compatibility with an enzymatic system as well as an excellent capability to be modified. The obtained biocomposite is also screen-printable, permitting the construction of glucose biosensors in planar configuration. A thick film device based on an optimized biocomposite showed a sensitivity (nA mM − 1) six times greater than other analogous

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materials and detection principles previously reported, which were used in the construction of a glucose biosensor based on a conventional amperometric electrode.

[19]

Acknowledgements [20]

We are grateful for the economic support of CICYT, Madrid (project BI096 – 0740). Carlos A. Gala´n-Vidal acknowledges a grant provided by DGAPA UNAM, Mexico.

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Biographies Carlos A. Gala´n-Vidal received the M.S. in 1993 from the Universidad Nacional Auto´noma de Me´xico. In 1994, he was awarded a grant from the Mexican government to work in the Sensor and Biosensor group of the Department of Chemistry of the Universitat Auto`noma de Barcelona. He received the Ph.D. degree in Analytical Chemistry (Thick-film biosensors) from the Universitat Auto`noma de Barcelona in 1996. In 1997, he joined Electrochemistry Unit from the Universidad Auto´noma de Hidalgo, where he is devoted to the development of planar biosensors. Ja6ier Mun˜oz received the Ph.D. degree in Physical Chemistry (Electrochemistry) from the Universitat Auto`noma de Barcelona in 1990. In 1990, he joined the Centro Nacional de Microelectro´nica of Barcelona (CSIC). In the same year, he also received a research fellowship from the Spanish Ministry of Science and Education to spend 2 years in the MESA Research Institute of the Universiteit Twente in the Prof P. Bergveld’s Laboratory. Since 1992, his main research interests include development of technological processes on silicon microelectronic technology for integrated optics and solid state chemical microsensors. Currently, he is working on the development of microsensors based on ISFET tranducers and amperometric biosensors, as well as in the development of UV curable polymers compatible with standard semiconductor processing. Carlos Domı´nguez received the Ph.D. degree in chemistry from the Universidad Complutense of Madrid in

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1985. He became a member of the scientific staff at the Centro Nacional de Microelectro´nica of Barcelona (CSIC) in 1986. Since 1991, he is Senior Scientific Researcher at CNM-CSIC. He is the head of the chemical microsensors group, where different research lines devoted to the development of chemical microsensors based on ISFET transducers and optoelectronics technology based on silicon for optochemical sensors, as well as the development of CVD and polymeric materials for applications in previous transducers.

.

Sal6ador Alegret became Professor of Analytical Chemistry from the Universitat Auto`noma de Barcelona in 1991. He is the head of the Sensor and Biosensor group of the Department of Chemistry of the UAB. He is currently devoted to the development of electrochemical biosensors based on enzymatic and immunological systems. The resulting sensors are being applied in automated analytical instrumentation for monitorization and process control on different fields, such as biomedicine, environment and chemical industry.