Preliminary investigations on a glucose biosensor based on the potentiometric principle

Sensors and Actuators B 123 (2007) 720–726

Preliminary investigations on a glucose biosensor based on the potentiometric principle Cheng-Wei Liao a , Jung-Chuan Chou b,∗ , Tai-Ping Sun c , Shen-Kan Hsiung d , Jui-Hsiang Hsieh a a

b

Institute of Biomedical Engineering, Chung Yuan Christian University, Chung-Li 320, Taiwan, ROC Graduate School of Electronic Engineering, National Yunlin University of Science and Technology, 123, Sec.3, University Rd., Douliou, Yunlin 640, Taiwan, ROC c Institute of Electrical Engineering, National Chi Nan University, Nantou 545, Taiwan, ROC d Institute of Electronic Engineering, Chung Yuan Christian University, Chung-Li 320, Taiwan, ROC Received 20 August 2006; received in revised form 5 October 2006; accepted 5 October 2006 Available online 7 November 2006

Abstract In this study, an electron mediator and a simple immobilization process are adopted to fabricate a potentiometric glucose biosensor. The enzyme and the electron mediator are immobilized on the surface of tin oxide (SnO2 )/indium tin oxide (ITO) glass using a covalent bond method to develop a disposable potentiometric glucose biosensor. The SnO2 /ITO glass is a pH sensor fabricated by depositing SnO2 thin films onto an ITO glass. The glucose oxidase (GOD) and the electron mediator (ferrocenecarboxylic acid, FcA) are co-immobilized on the SnO2 /ITO glass using 3-glycidyloxypropyltrimethoxysilane (GPTS). This work investigates the coimmobilization of GOD and FcA as a useful approach for enlarging the dynamic range to a glucose concentration of 360 mg/dl, and for improving linearity and sensitivity of the fabricated glucose biosensor. The experimental results indicate that the optimal weight ratio of GOD to FcA is 1:1. The output signal is associated with the pH of the measurement environment and the optimal pH value is pH 7.5. © 2006 Elsevier B.V. All rights reserved. Keywords: Glucose oxidase (GOD); 3-Glycidyloxypropyltrimethoxysilane (GPTS); Tin oxide (SnO2 ); Electron mediator; Potentiometric biosensor

1. Introduction Diabetes is still a metabolic disorder that is caused by a total or partial lack of insulin. Glucose fluctuations within the normal physiological range of 110 ± 25 mg/dl are considered to be acceptable; diabetics had values of 360 mg/dl or higher [1]. For almost four decades, researchers engaged in the development of glucose-sensing devices which monitored the glucose levels in biological fluids rapidly, accurately and continuously, especially to help type II diabetes mellitus patients to monitor their daily sugar levels [2]. Spectrophotometric approaches were laboratory methods, which were not useful in on-line monitoring. The inconvenience of spectrophotometric methods was overcome using analyzers based on electrochemical methods, in which biosensors were applied. Numerous biosensors were employed to

∗

Corresponding author. Tel.: +886 5 5342601x2101; fax: +886 5 5321719. E-mail address: [email protected] (J.-C. Chou).

0925-4005/$ – see front matter © 2006 Elsevier B.V. All rights reserved. doi:10.1016/j.snb.2006.10.006

measure glucose levels in biological samples. Among these biosensors, amperometric glucose biosensors were attracted significant interest. However, applying a high polarizing voltage (Eapp = 0.6–0.8 V) oxidized interfering substances such as ascorbic acid and uric acid, which were commonly present in biological fluids, leading to nonspecific signals [3]. Several artificial redox mediators were investigated as electron acceptors, to reduce the applied potential in an amperometric glucose biosensor, and thereby solved this problem. Electron mediators were incorporated to reduce the potential applied to the working electrode below 0.8 V. Hence, some investigations demonstrated the effects of applying electron mediators [2,4–12]. Furthermore, base electrodes were modified to improve the performance of glucose biosensors [13–17]. Accordingly, in glucose biosensors based on the amperometric principle, polarizing voltage is the key factor leading to interference. Since no extra potential was required to apply on potentiometric biosensors, interference caused by the polarizing voltage could be eliminated. The enzyme field effect transistor (EnFET), based on the ion-sensitive field effect transistor (ISFET), was

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first introduced by Caras and Janata [18]. Glucose oxidase typically hydrolyzes glucose according to the equations shown as follows [19]: ␤-d-Glucose + O2

␤-d-glucose oxidase

−→

d-glucono-␦-lactone + H2 O2 (1)

d-Glucono-δ-lactone → d-gluconate + H+

(2)

In that system, the common electron acceptor oxygen (O2 ) produced the product hydrogen peroxide (H2 O2 ). ISFETs measured the glucose concentration by detecting the variation in pH caused by the generation of hydrogen ions by the dissociation of glucose acid. Due to the low dissociation constant (pKa ∼ = 3.8) of glucose acid, ISFET glucose sensors were responsible for their low sensitivities [20]. The glucose concentration in human blood was typically around 110 mg/dl, reaching 360 mg/dl or more for diabetics. However, the concentration of oxygen in arteries and veins did not exceed 14.5 and 5.85 kPa, respectively. Since the concentration ratio of oxygen in real blood was unfavorable, the dynamic range of the biosensor was normally limited by the oxygen and did not exceed several mM. Additionally, hydrogen peroxide, one of the by-products of glucose oxidation, inhibited glucose oxidase, reducing the sensitivity and repeatability of a steady glucose biosensor measurement system [21]. Seo et al. [21] and Lee et al. [22] utilized a platinum (Pt) electrode actuator on an ISFET sensitive gate to electrolyze the hydrogen peroxide. Their sensors, coupled with the Pt electrode actuator, exhibited a wide dynamic range, from 18 to 180 mg/dl. Yin et al. [19] developed an amorphous tin oxide (SnO2 )/indium tin oxide (ITO) glass structure with a sensitive gate, coimmobilized with glucose oxidase and manganese dioxide (MnO2 ) by a cross-linking method, they developed the sensor as a disposable glucose EnFET. MnO2 was employed as a catalyst, as it catalyzed the decomposition of hydrogen peroxide to H2 O and O2 . The glucose EnFET doped with MnO2 was found to be useful in extending the dynamic range up to a glucose concentration of 360 mg/dl and the output signal was enlarged to 50 mV. The technology used to process ISFET was strongly related to the various MOSFET technologies [23]. The dimension of MOSFET was reduced to the nanoscale, affecting the size of the sensing area, and thereby, the sensitivity. For reasons of sensitivity, the sensing areas of most biosensors were limited to the micro scale [24]. The ISFET fabrication processes were incompatible with the fast-developing MOSFET technologies. This incompatibility represented an obstacle to the development of ISFEF. In our study, the SnO2 /ITO glass pH sensor was fabricated by depositing SnO2 thin films onto ITO glass, which was not limited by MOSFET technologies. This approach had numerous advantages, including high sensitivity, ease of fabrication and low cost [25,26]. The use of a low-cost substrate and the simple fabrication process reduced the cost of the potentiometric glucose biosensor made from an ITO substrate below that made from an ISFET substrate. To our knowledge, potentiometric glucose biosensors were based on the principles of the electrolyzation

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[21,22] or catalyzation [19] of H2 O2 , to improve their performance. An artificial redox electron mediator was used for the first time herein to develop a potentiometric glucose biosensor and thereby solve the problems of O2 deficiency, and the H2 O2 inhibition effect on glucose oxidases (GOD) problems. Ferrocene and its derivatives were some of the most efficient electron mediators [2,5]. Among the ferrocene derivatives, ferrocenecarboxylic acid (FcA) was the optimal alternative to be used in the fabrication, because of its hydrophilic characteristic, so it could be mixed with the enzyme in a single step without any pretreatment [27–29]. Moreover, we discovered that 3-glycidyloxypropyltrimeth-oxysilane (GPTS) [30,31] acted as a very effective and easily applied attachment medium to coimmobilize FcA and GOD on the SnO2 /ITO glass pH sensor. In this study, a potentiometric glucose biosensor based on the SnO2 /ITO glass pH sensor was realized, by applying GPTS to coimmobilize enzyme and electron mediator on its surface. The following equations describe the reaction sequences [10]: Glucose + GOD(FAD) → Gluconolactone + GOD(FADH2 ) (3) GOD(FADH2 ) + 2FcA(ox) → GOD(FAD) + 2FcA(red) + 2H+ (4) Herein, ferrocenecarboxylic acid (FcA) is employed as an electron mediator. In the following reaction, the active center flavin adenine dinucleotide (FAD) of glucose oxidases oxidizes the glucose penetrating the membrane; electrons are then transferred from the reduced GOD(FADH2 ) to FcA(ox) . The resulting product, H+ , is detected by the SnO2 /ITO glass pH sensor. Accordingly, applying this mediator to a potentiometric glucose biosensor and using GPTS to simplify the immobilization process was expected to produce excellent response characteristics when the biosensor was exposed to glucose in a physiological environment. Furthermore, because no extra potential was required to apply on the potentiometric glucose biosensor, other biological molecules presented in physiological fluids will not be capable of reacting at the electrode surface. 2. Experiment 2.1. Chemicals and materials Glucose oxidase (GOD; EC 1.1.3.4, type VII, from aspergillus niger) with an activity value of 121 units/mg was purchased from the Sigma Company. 3-Glycidyloxypropyltrimethoxysilane (GPTS) and ferrocenecarboxylic acid (99%) were obtained from the Aldrich Company. All other reagents were of reagent grade and were used without further purification. Deionized (D.I.) water was used to make all of the electrolytes and the buffer solutions. Tin oxide thin films were formed using a radio frequency sputtering system (tin oxide target, 99.9%) at a substrate temperature of 150 ◦ C. ITO glass (50–100 /; ITO coating thickness 23 nm) was supplied by the Wintek Corporation.

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Fig. 1. Cross-section of sensing structure for the SnO2 /ITO glass pH sensor with enzyme biomembrane.

2.2. Sensor fabrication In this work, the potentiometric glucose biosensor was based on a SnO2 /ITO glass pH sensor. Fig. 1 depicts the sensitive part of the SnO2 /ITO glass pH sensor with an enzyme membrane. The SnO2 thin films were deposited by the sputtering method with a thickness of 200 nm and a sensing area of 2 mm × 2 mm. Before the SnO2 thin films were deposited on the ITO glass, the ITO glass was cleaned in methyl alcohol solution and D.I. water for 20 and 10 min, respectively. The SnO2 /ITO glass pH sensor exhibited a linear pH response of around 58mV/pH between pH 2.0 and 12.0 [25]. 2.3. Enzyme immobilization The processes for coating GPTS are as follows: (a) the SnO2 /ITO glass pH sensor was cleaned with D.I. water for 15 min; (b) 2 ␮l of the GPTS was dropped onto the sensing window of the SnO2 /ITO glass pH sensor; (c) the SnO2 /ITO glass pH sensor coated with the GPTS was cured at 150 ◦ C for 2 h; (d) following curing, the device was washed by immersing in 5 mM, pH 7 phosphate buffer solution (PBS) for 10 min, to clean away the GPTS that was not bound to the sensing window. In this work, a GPTS was adopted to immobilize the enzyme and electron mediator on the SnO2 /ITO glass pH sensor [30,31]. The GTPS was dropped onto the SnO2 thin films to form strong covalent bonds with the surface of these films. Finally, the mixture of enzyme and electron mediator was dropped onto the GPTS membrane. The GPTS contained hydrolizable silicon alkoxide bonds and a ring-opening epoxy group. Accordingly, the enzyme and electron mediator were immobilized on the SnO2 /ITO glass pH sensor in a single step. Therefore, the GPTS coimmobilized the enzyme and mediator on the SnO2 /ITO glass pH sensor. After the reactions of enzyme, electron mediator and glucose, the resulting product, H+ , was detected by the SnO2 /ITO glass pH sensor. In the preparation of the enzyme membrane, a liquid solution was obtained by mixing the GOD and FcA in 5 mM, pH 7.5 phosphate buffer solution. Then, a 1.5 ␮l aliquot of the mixture was dropped onto the GPTS membrane, which was employed to immobilize the enzyme and electron mediator. Finally, the sensors were shielded from light and allowed to soak for 12 h at 4 ◦ C.

Fig. 2. Measurement circuit with an instrumentation amplifier LT1167.

2.4. Measurement system The readout circuit was based on an instrumentation amplifier, LT1167, shown in Fig. 2, which was utilized to measure the total charges associated with the hydrogen ions accumulated on the sensing window. Where the amplifier gain in Fig. 2 was fixed at unity, the measurement configuration consisted of a SnO2 /ITO glass pH sensor with a biomembrane and an Ag/AgCl reference electrode. The instrumentation amplifier, LT1167, was configured as a linear, small signal amplifier whose input voltage was connected to the pH potentiometric transducer. All the measurements in this study were made at room temperature. 3. Results and discussion 3.1. Response time of the potentiometric glucose biosensor In this study, a separately sensitive gate biomembrane/ SnO2 /ITO glass structure was employed as a disposable biochemical transducer. This structure provided the advantages of insensitivity to light, simplicity of fabrication and a lower cost than the traditional ISFET, the silicon-on-sapphire (SOS) structure ISFET or the silicon-based EGFET [23]. Fig. 3 plots the pH response of the sensitive structure. It showed a sensitivity of 58.13 mV/pH at pH values of between 2.03 and 11.53, which was obtained from the slope of Fig. 3. The potentiometric glucose biosensor was immersed into blank buffer solution and then in 360 mg/dl glucose solution. Fig. 4 shows the results. The reaction curve exhibited exponential dependence on time. When the potentiometric glucose biosensor was immersed in the blank buffer solution, no glucose was introduced into it. Since no signal was produced at the output of the amplifier, the voltage response curve formed a straight line, which is the baseline in Fig. 4. The results showed that the voltage responses at the 50th, 230th and 600th second are −0.16145, −0.0601 and

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Fig. 3. Output voltage vs. pH of the SnO2 /ITO glass pH sensor connected to an instrumentation amplifier LT1167.

−0.05405 V, respectively. Since the voltage response of the potentiometric glucose biosensor immersed in blank buffer solution for 50 s had been subtracted, the output voltages (V) at the 230th and 600th second were 101.05 and 107.4 mV, respectively. Thus, the output voltage reached 95% within 3 min. The glucose EnFETs developed by Yin et al. [19], without and with an outer bovine serum albumin (BSA) membrane, had response time of 5 and 12 min, respectively. In this study, the GPTS was employed to coimmobilize the GOD and FcA in a single step, simplifying the immobilization. Yin et al. [19] applied a cross-linking approach to immobilize enzyme on the sensing gate and then a BSA membrane doped with MnO2 was dropped onto the enzyme membrane. Clearly, the immobilization method and membrane thickness affected the response time, because the generated hydrogen ions reached the sensitive gate region through the biomembrane. Therefore, a simple immobilization procedure and a thin membrane accelerated the diffusion of the hydrogen ions, improving the response time.

Fig. 4. Voltage response of the potentiometric glucose biosensor used to detect 360 mg/dl glucose in 5 mM, pH 7.5 buffer solution.

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Fig. 5. Calibration curve of the potentiometric glucose biosensor. Sensors were measured in 5 mM, pH 7.5 buffer solution.

3.2. Calibration curve of glucose biosensor Fig. 5 plots the calibration curve of the potentiometric glucose biosensor. Each glucose concentration was measured eight times from eight sensors. Calculations produced the reliability of the potentiometric glucose biosensor. The potentiometric glucose biosensor incorporated with FcA exhibited a strong response magnitude and high sensitivity. The sensitivity of the potentiometric glucose biosensor developed herein was 0.256 mV(mg/dl)−1 , which exceeded that of the glucose EnFET [19], 0.094 mV(mg/dl)−1 . Two factors were believed to be responsible for these improvements. The first is that the GPTS did not eliminate the activity of the GOD, and the other is that FcA replaced oxygen as the electron acceptor. 3.3. Effect of weight ratio of GOD to FcA ISFET-based glucose biosensors based on the potentiometric principle detected gluconic acid produced by the enzymatic conversion of glucose molecules. However, the by-product of this reaction was hydrogen peroxide, which degraded the enzyme [23]. Hence, hydrogen peroxide constituted most of the noise for the potentiometric glucose biosensor. In this study, an electron mediator was introduced into the glucose reactions. Fig. 6 displays the potentiometric glucose biosensor responses for sensors, incorporated with and without FcA. The results demonstrated that sensors incorporated with FcA showed a wider dynamic range and better linearity than those without FcA, suggesting that the potentiometric glucose biosensor incorporated with FcA overcame the problems of O2 deficiency and the H2 O2 inhibition effect on GOD, even at high glucose concentrations, since the electron mediator acted as an electron acceptor and also catalyzed the decomposition of H2 O2 to H2 O and O2 [3,10]. The function of the electron mediator, which catalyzed the decomposition of hydrogen peroxide, caused the potentiometric glucose biosensor to eliminate the hydrogen peroxide. Therefore, the degradation by hydrogen peroxide deterioration of GOD was neglected.

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Fig. 6. Voltage response of the potentiometric glucose biosensor, incorporated with and without FcA.

Fig. 7. Calibration curves of the potentiometric glucose biosensors at various weights of FcA. The weight of GOD was fixed at 1 mg.

In this study, the FcA and GOD were coimmobilized on the surface of SnO2 thin films by GPTS in a single step. The function of glucose oxidase was utilized to oxidize glucose. Meanwhile, the mediator acted as an electron acceptor and also catalyzed the decomposition of H2 O2 to H2 O and O2 [3,10]. The percentages of GOD and FcA in the biomembrane affected the performance of glucose oxidation, the acceptance of electrons and the decomposition of H2 O2 to H2 O and O2 . Accordingly, the other effects of the electron mediators, overcoming O2 deficiency and the H2 O2 inhibition effect on GOD, were verified by the potentiometric principles herein. Although electron mediators were commonly adopted in fabricating amperometric glucose biosensors, only a few investigations established the optimal weight ratio of GOD and electron mediator. Liu et al. [7] and Parellada et al. [32] showed that the optimal weight ratio of GOD to the electron mediator was 1:1. In 2005, Fei et al. [12] reported a glucose nanosensor based on redox polymer/glucose oxidasemodified carbon fiber nanoelectrode, whose optimal weight ratio of GOD to the electron mediator was 2:3. The volume of the buffer solution applied to mix GOD and FcA was 100 ␮l. Then, the 1.5 ␮l mixture was dropped onto the sensing window with an area of 2 mm × 2 mm. The mass of the GOD was fixed at 1 mg; the mass of FcA was then adjusted to 0.5, 1, 2, 2.5 and 3 mg, respectively. The experimental results shown in Fig. 7 demonstrate that the linearities at FcA = 0.5, 1, 2, 2.5 and 3 mg are 0.955, 0.995, 0.989, 0.977 and 0.981, respectively. The preferred weight value of FcA was 1 mg, which was determined by the linearity. FcA and GOD masses of 1 mg each were preferred. Therefore, the optimal weight ratio of GOD to FcA for the potentiometric glucose biosensor was 1:1, because of the enhanced efficiencies of glucose oxidation, electron transfer and decomposition of H2 O2 to H2 O and O2 between GOD and FcA at this weight ratio.

mance of the potentiometric glucose biosensor. The key factors in determining the efficiencies of glucose oxidation, electron transfer and decomposition of H2 O2 were the enzyme activity and capacity of the electron mediator. According to the experimental results presented in Fig. 8, the linearities at pH 7, 7.5, 8 and 8.5 are 0.872, 0.995, 0.971 and 0.968, respectively. The activity of the GOD was reduced in alkaline environments [33,34], suggesting that the optimal environment was acid solution. However, the potentiometric glucose biosensor exhibited low reproducibility in pH 7 solution, revealing that the electron mediator could not strongly decompose H2 O2 in an acid environment [19]. As the electron mediator weakly decomposed H2 O2 , H2 O2 inhibited the activity of the enzyme [21], resulting in low reproducibility in pH 7 solution. The enzyme activity and the ability of the electron mediator to decompose H2 O2 were such that the potentiometric glucose biosensor immersed in pH 7.5 solution exhibited high linearity and reproducibility. Therefore, pH 7.5 solution was selected as the optimal environment herein.

3.4. Influence of pH The efficiencies of glucose oxidation, electron transfer and the decomposition of H2 O2 to H2 O and O2 affected the perfor-

Fig. 8. Calibration curves of the potentiometric glucose biosensor in buffer solutions at various pH values.

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4. Conclusion In this study, the electron mediator was successfully applied to the potentiometric glucose biosensor based on the SnO2 /ITO glass pH sensor. Additionally, the GPTS was employed to coimmobilize the enzyme and the electron mediator in a single step, simplifying the immobilization processes. A disposable potentiometric glucose biosensor with these advantages was realized herein. The cost of the separative sensitive gate potentiometric glucose biosensor was less than that of a traditional ISFET. Furthermore, the mixture of the enzyme and electron mediator was directly coimmobilized on the SnO2 thin films via covalent linkage using GPTS, which was fabricated more easily than by traditional immobilization methods, such as crosslinking and gel entrapment. In this study, FcA replaced O2 as an electron acceptor in the catalytic decomposition of H2 O2 . The overcoming of the O2 deficiency and the H2 O2 inhibition effect on GOD enabled the potentiometric glucose biosensor to exhibit good linearity, high sensitivity and a wide dynamic range up to a glucose concentration of 360 mg/dl. The optimal weight ratio of GOD to FcA was 1:1, based on the linearity, because of the enhanced efficiencies of glucose oxidation, electron transfer and the decomposition of H2 O2 to H2 O and O2 between the GOD and FcA at this weight ratio. Moreover, the enzyme activity and the ability of the electron mediator to decompose H2 O2 were such that the potentiometric glucose biosensor immersed in pH 7.5 solution exhibited high linearity and reproducibility. Hence, pH 7.5 solution was chosen as the optimal environment herein. Since the pH value of human blood was around 7.4, the potentiometric glucose biosensor was suitable for measuring the concentration of glucose of human blood. Acknowledgement

[8]

[9]

[10]

[11]

[12]

[13]

[14]

[15]

[16]

[17] [18] [19]

[20]

The authors would like to thank the National Science Council of the Republic of China, Taiwan for financially supporting this research under Contract Nos. NSC 94-2213-E-033-005 and NSC 95-2221-E-033-028.

[21]

References

[23]

[1] E. Wilkins, P. Atanasov, Glucose monitoring: state of the art and future possibilities, Med. Eng. Phys. 18 (4) (1996) 273–288. [2] H. Patel, X. Li, H.I. Karan, Amperometric glucose sensors based on ferrocene containing polymeric electron transfer systems–a preliminary report, Biosens. Bioelectron. 18 (2003) 1073–1076. [3] C. Mousty, Sensors and biosensors based on clay-modified electrode—new trends, Appl. Clay Sci. 27 (2004) 159–177. [4] R. Nagata, S.A. Clark, K. Yokoyama, E. Tamiya, I. Karube, Amperometric glucose biosensor manufactured by a printing technique, Anal. Chim. Acta 304 (1995) 157–164. [5] G.W.J. Harwood, C.W. Pouton, Amperometric enzyme biosensors for the analysis of drugs and metabolites, Adv. Drug Deliv. Rev. 18 (1996) 163–191. [6] T. Saito, M. Watanabe, Characterization of poly(vinylferrocene-co-2hydroxyethyl methacrylate) for use as electron mediator in enzymatic glucose sensor, React. Funct. Polym. 37 (1998) 263–269. [7] H. Liu, H. Li, T. Ying, K. Sun, Y. Qin, D. Qi, Amperometric biosensor sensitive to glucose and lactose based on co-immobilization of ferrocene,

[22]

[24]

[25]

[26]

[27]

[28]

[29]

725

glucose oxidase, ␤-galactosidase and mutarotase in ␤-cyclodextrin polymer, Anal. Chim. Acta 358 (1998) 137–144. F. Ge, X.E. Zhang, Z.P. Zhang, X.M. Zhang, Simultaneous determination of maltose and glucose using a screen-printed electrode system, Biosens. Bioelectron. 13 (1998) 333–339. X. Yang, L. Hua, H. Gong, S.N. Tan, Covalent immobilization of an enzyme (glucose oxidase) onto a carbon sol–gel silicate composite surface as a biosensing platform, Anal. Chim. Acta 478 (2003) 67–75. S. Berchmans, R. Sathyajith, V. Yegnaraman, Layer-by-layer assembly of 1,4-diaminoanthraquinone and glucose oxidase, Mater. Chem. Phys. 77 (2002) 390–396. B. Alonso, P.G. Armada, J. Losada, I. Cuadrado, B. Gonz´alez, C.M. Casado, Amperometric enzyme electrodes for aerobic and anaerobic glucose monitoring prepared by glucose oxidase immobilized in mixed ferrocenecobaltocenium dendrimers, Biosens. Bioelectron. 19 (2004) 1617–1625. J. Fei, K. Wu, F. Wang, S. Hu, Glucose nanosensors based on redox polymer/glucose oxidase modified carbon fiber nanoelectrodes, Talanta 65 (2005) 918–924. D. Pan, J. Chen, S. Yao, L. Nie, J. Xia, W. Tao, Amperometric glucose biosensor based on immobilization of glucose oxidase in electropolymerized o-aminophenol film at copper-modified gold electrode, Sens. Actuators B 104 (2005) 68–74. W.J. Guan, Y. Li, Y.Q. Chen, X.B. Zhang, G.Q. Hu, Glucose biosensor based on multi-wall carbon nanotubes and screen printed carbon electrodes, Biosens. Bioelectron. 21 (2005) 508–512. S. Zhang, N. Wang, H. Yu, Y. Niu, C. Sun, Covalent attachment of glucose oxidase to an Au electrode modified with gold nanoparticles for use as glucose biosensor, Bioelectrochemistry 67 (2005) 15–22. X. Cui, G. Liu, Y. Lin, Amperometric biosensors based on carbon paste electrodes modified with nanostructure mixed-valence manganese oxides and glucose oxidase, Nanomed. Nanotechnol. Biol. Med. 1 (2005) 130–135. H. Yang, Y. Zhu, Size dependence of SiO2 particles enhanced glucose biosensor, Talanta 68 (2006) 569–574. S. Caras, J. Janata, Field effect transistors sensitive to penicillin, Anal. Chim. Acta 52 (1980) 1935–1937. L.T. Yin, J.C. Chou, W.Y. Chung, T.P. Sun, K.P. Hsiung, S.K. Hsiung, Glucose biosensor doped with MnO2 powder, Sens. Actuators B 76 (2001) 187–192. A.A. Shul’ga, A.C. Sandrovsky, V.I. Strikha, A.P. Soldatkin, N.F. Starodub, A.V. El’skaya, Overall characterization of ISFET-based glucose biosensor, Sens. Actuators B 10 (1992) 41–46. H.I. Seo, C.S. Kim, B.K. Sohn, T. Yeow, M.T. Son, M. Haskard, ISFET glucose sensor based on a new principle using the electrolysis of hydrogen peroxide, Sens. Actuators B 40 (1997) 1–5. C.H. Lee, H.I. Seo, Y.C. Lee, B.W. Cho, H. Jeong, B.K. Sohn, All solid type ISFET glucose sensor with fast response and high sensitivity characteristics, Sens. Actuators B 64 (2000) 37–41. P. Bergveld, Thirty years of ISFETOLOGY, Sens. Actuators B 88 (2003) 1–20. R.A. Yotter, D.M. Wilson, Sensor technologies for monitoring metabolic activity in single cells—Part II: nonoptical methods and applications, IEEE Sens. J. 4 (4) (2004) 412–429. L.T. Yin, J.C. Chou, W.Y. Chung, T.P. Sun, S.K. Hsiung, Separate structure extended gate H+ -ion sensitive field effect transistor on a glass substrate, Sens. Actuators B 71 (2000) 106–111. L.T. Yin, J.C. Chou, W.Y. Chung, T.P. Sun, S.K. Hsiung, Study of indium tin oxide thin film for separative extended gate ISFET, Mater. Chem. Phys. 70 (2001) 12–16. R. Nagata, S.A. Clark, K. Yokoyama, E. Tamiya, I. Karube, Amperometric glucose biosensor manufactured by a printing technique, Anal. Chim. Acta 304 (1995) 157–164. O. Elekes, D. Moscone, K. Venema, J. Korf, Bi-enzyme reactor for electrochemical detection of low concentrations of uric acid and glucose, Clin. Chim. Acta 239 (1995) 153–165. N. Ganesan, A.P. Gadre, M. Paranjape, J.F. Currie, Gold layer-based dual cross-linking procedure of glucose oxidase with ferrocene monocarboxylic acid provides a stable biosensor, Anal. Biochem. 343 (2005) 188–191.

726

C.-W. Liao et al. / Sensors and Actuators B 123 (2007) 720–726

[30] Y. Park, M. Nagai, Proton exchange nanocomposite membranes based on 3-glycidoxypropyltrimethoxysilane, silicotungstic acid and ␣-zirconium phosphate hydrate, Solid State Ionics 145 (2001) 149–160. [31] H. Zhu, M. Snyder, Protein arrays and microarrays, Curr. Opin. Chem. Biol. 5 (2001) 40–45. [32] J. Parellada, A. Narvaez, E. Dominguez, I. Katakis, A new type of hydrophilic carbon paste electrodes for biosensor manufacturing: binder paste electrodes, Biosens. Bioelectron. 12 (4) (1997) 267–275. [33] T. Godjevargova, V. Konsulov, A. Dimov, Preparation of an ultrafiltration membrane from the copolymer of acrylonitrile-glycidylmethacrylate utilized for immobilization of glucose oxidase, J. Membr. Sci. 152 (1999) 235–240. [34] V. Bulmus, H. Ayhan, E. Piskin, Modified PMMA monosize microbeads for glucose oxidase immobilization, Chem. Eng. J. 65 (1997) 71–76.

Biographies Cheng-Wei Liao was born in Nantou, Taiwan, Republic of China, on December 10, 1964. He received the BS degree in biomedical engineering from Chung Yuan Christian University, Chung-Li, Taiwan, in 1987; the MS degree in biomedical engineering from Chung Yuan Christian University, Chung-Li, Taiwan, in 1992. Since 1993 he has been patent examiner at the Intellectual Property Office, Ministry of Economic Affairs. And since 2002, he has been working toward the PhD degree in the Department of Biomedical Engineering at Chung Yuan Christian University, Chung-Li, Taiwan. His research interests include the biosensor and its applications. Jung-Chuan Chou was born in Tainan, Taiwan, Republic of China, on July 13, 1954. He received the BS degree in physics from Kaohisung Normal College, Kaohsiung, Taiwan, in 1976; the MS degree in applied physics from Chung Yuan Christian University, ChungLi, Taiwan, in 1979; the PhD degree in electronics from National Chiao Tung University, Hsinchu, Taiwan, in 1988. He taught at Chung Yuan Christian University from 1979 to 1991. Since 1991 he has worked as an associate professor in the Department of Electronic Engineering at the National Yunlin University of Science and Technology. From 1997 to 2002, he was Dean, Office of Technology Cooperation at the National Yunlin University of Science and Technology. And since 2002, he has been Chief Secretary at the National Yunlin University of Science and Technology. His research interests are in the areas of amorphous materials and devices, electrographic photoreceptor materials and devices, electronic materials and devices, sensor devices and science education.

Tai-Ping Sun was born in Taiwan on March 20, 1950. He received the BS degree in electrical engineering from Chung Cheng Institute of Technology, Taiwan, in 1974, the MS degree in material science engineering from National Tsing Hua University, Taiwan, in 1977, and PhD degree in electrical engineering from National Taiwan University, Taiwan, in 1990. From 1977 to 1997, he worked at Institute of Science and Technology, Republic of China, concerning the development of Infrared device, circuit and system. He joined the Department of Management Information System, Chung-Yu College of Business Administration since 1997 as an associate professor. Since 1999 he has joined the Department of Electrical Engineering, National Chi Nan University as a professor. And since 2001, he has been Secretary-General at the National Chi Nan University. And his research interests are infrared detector and system, analog/digital mixed-mode integrated circuit design, special semiconductor sensor and their applications. Shen-Kan Hsiung was born on June 14, 1942. He received the BS degree from Department of Electrical Engineering, National Cheng-Kung University, in 1965; the MS degree from Department of Electronic Engineering, National Chiao-Tung University, Taiwan, in 1968; the PhD degree from Material Science Engineering of USC, USA, in 1974. From 1974 to 1978, he was an associate professor in Department of Electrical Engineering, Chung Yuan Christian University. Since 1978 he has been a professor in Department of Electronic Engineering, Chung Yuan Christian University. And since 2000, he has been President, at the Chung Yuan Christian University. His current interests are electronic materials, amorphous thin films and semiconductor sensors. Jui-Hsiang Hsieh was born on September 8, 1953. She received the BS degree from the Department of Biology, National Taiwan Normal University, Taiwan, in 1979, the MS degree from the Institute of Biology, National Taiwan Normal University, Taiwan, in 1982 and the PhD degree from the Institute of Zoology, National Taiwan University, Taiwan, in 1995. From 1989 to 1999, she was an associate professor in Department of Biomedical Engineering, Chung Yuan Christian University. Since 1999 she has been a professor in the Department of Biomedical Engineering, Chung Yuan Christian University. Her current interests are in the study of the correlationship between the cardiovascular areas and the respiration of the lower brain stem in cat or turtle.