Glucose biosensing based on the highly efficient immobilization of glucose oxidase on a Prussian blue modified nanostructured Au surface

Electrochemistry Communications 11 (2009) 2048–2051

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Glucose biosensing based on the highly efficient immobilization of glucose oxidase on a Prussian blue modified nanostructured Au surface Asieh Ahmadalinezhad, A.K.M. Kafi, Aicheng Chen * Department of Chemistry, Lakehead University, Thunder Bay, Ontario, Canada P7B 5E1

a r t i c l e

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Article history: Received 13 July 2009 Received in revised form 23 August 2009 Accepted 25 August 2009 Available online 29 August 2009 Keywords: Glucose Nanostructured Au Glucose oxidase Biosensor Prussian blue

a b s t r a c t We report on a novel glucose biosensor based on the immobilization of glucose oxidase (GOx) on a Prussian blue modified nanoporous gold surface. The amperometric glucose biosensor fabricated in this study exhibits a fast response and the very low detection limit of 2.5 lM glucose. The sensitivity of the biosensor was found to be very high, 177 lA/mM; the apparent Michaelis–Menten constant is calculated to be 2.1 mM. In addition, the biosensor has good reproducibility and remains stable over 60 days. The anti-interference ability of the biosensor was also assessed, showing little interference from possible interferents such as ascorbic acid (AA), acetaminophen (AP) and uric acid (UA). Ó 2009 Elsevier B.V. All rights reserved.

1. Introduction Certain areas of medical, food and environmental analysis require express and inexpensive monitoring methods, e.g. for glucose in blood and food [1,2]. There are many different methods proposed to measure glucose concentrations such as fluorescence spectroscopy [3], near infrared spectroscopy [4], Raman spectroscopy [5], nuclear magnetic resonance spectroscopy [6], liquid chromatography–mass spectroscopy [7], high performance liquid chromatography [8], capillary electrophoresis [9], polarography [10], and electrochemical methods [11–15]. In the area of electrochemical methods, biosensors represent an important subclass of sensors in which an electrode is used as the transduction element [16,17]. Enzymatic biosensor performance strongly depends on the electron transfer between the enzyme and electrode. Proper immobilization of the enzyme on the electrode surface is critical to enzymatic biosensor development. It has been shown that gold nanoparticles are able to enhance the electron transfer process [18]. In addition, it is known that Prussian blue (ferric hexacyanoferrate) is an efficient redox mediator for the selective detection of H2O2 in the presence of oxygen and other interferents [19,20]. Glucose oxidase, as an enzyme, catalyses the oxidation of b-D-glucose to D-glucono-1,5-lactone and hydrogen peroxide using molecular oxygen as the electron acceptor. In this communication, for the first time, we report on the direct growth of nanoporous Au on Ti substrates and the immobilization of glucose oxidase on Prussian * Corresponding author. Tel.: +1 807 343 8318; fax: +1 807 346 7775. E-mail address: [email protected] (A. Chen). 1388-2481/$ - see front matter Ó 2009 Elsevier B.V. All rights reserved. doi:10.1016/j.elecom.2009.08.048

blue modified nanoporous Au surfaces. Our electrochemical study shows that the fabricated biosensor displays high sensitivity, a very low detection limit, high stability and long lifetime towards the detection of glucose. 2. Experimental 2.1. Materials Glucose oxidase (EC 1.1.3.4, Type VII from Aspergillus niger) and were purchased from Sigma and used as received. Chitosan was purchased from Aldrich, and HAuCl4 (99.9%) from Alfa Aesar. Ferric chloride (FeCl3) was purchased from Anachemia, and potassium ferricyanide (III) (K3Fe(CN)6) and ammonium formate (99.995%) from Sigma–Aldrich. All other reagents were of analytical grade. Water was purified with the NanopureÒ water system (18 MX cm) and was used to prepare all solutions and clean all materials. All experiments were performed in a 0.1 M phosphate buffer solution (PBS) with different pH values maintained with K2HPO4 and KH2PO4. D-glucose

2.2. Glucose biosensor fabrication The Ti substrate plates (1.25 cm  0.80 cm  0.5 mm) were first degreased in an ultrasonic bath of acetone for 10 min followed by 10 min in pure water (18 MX). The substrates were then etched in 18% HCl at approximately 85 °C for 30 min. The etched Ti substrates were transferred into Teflon-lined autoclaves and 10 ml of

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2.3. Biosensor characterization and electrochemical measurements The morphology and composition of the nanoporous Au materials were characterized by scanning electron microscope (SEM) (JEOL JSM 5900 LV) at an acceleration voltage of 13 kV, and an energy-dispersive X-ray (EDS) spectrometer. The X-ray diffraction (XRD) patterns of the as-prepared samples were recorded using a XRD Philips PW 1050-3710 diffractometer with Cu Ka radiation. The cyclic voltammograms (CVs) were acquired from 0.2 V to +0.3 V at a scan rate of 10 mV/s in 50 ml of 0.1 M pH 7.0 PBS. The amperometric response of the biosensor to glucose was recorded in a stirred PBS at 0.1 V vs. the Ag/AgCl reference electrode. All electrochemical experiments were performed at room temperature, 20 ± 2 °C. 3. Results and discussion 3.1. Characterization of the Ti/Au electrode surface Fig. 1a and b present a typical SEM image and EDS spectrum of the Ti/NPAu electrode, respectively, fabricated using the hydrothermal method. The SEM image reveals that nanoporous Au networks were formed and completely covered the substrate. The three-dimensional random porous structures, with diameters of tens to hundreds of nanometers, were formed by the close connection of the Au nanoparticles, with a size range of 50–500 nm. Only Au peaks appear in the EDS spectrum, and no discernible carbon or oxygen signals are seen in the figure, showing that the synthesized Au-nanoporous networks are free of surface organic impurities. 3.2. Electrochemical behaviour of the PB modified Au electrodes Fig. 2A displays the cyclic voltammograms of the PB modified Au electrodes recorded in a 0.1 M PBS (pH = 7.0) at a scan rate of 10 mV/s. A pair of well-defined oxidation and reduction peaks are observed for the Ti/Au/PB electrode (Plot a). The anodic and cathodic peak potentials are located at +0.052 V and 0.016 V, respectively. The formal potential E0 was calculated to be +0.018 V. The redox peaks significantly increase at the Ti/NPAu/ PB electrode (Plot b), showing that the nanoporous Au directly grown on the Ti substrate hydrothermally possesses a much higher

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an aqueous mixture of inorganic metal precursor and a reducing agent was added. The 1 M ammonium formate was used as the reducing agent and the metal precursor was 10 mM HAuCl4. The autoclaves were sealed and heated at 180 °C for 10 h. The Ti plates coated with Au were dried and annealed under argon at 250 °C for 2 h. After cooling to room temperature, Prussian blue (PB) was deposited onto the nanoporous Au surface by electrodeposition. This was accomplished by immersing the nanoporous Au (NPAu) electrode in a solution containing 2.5 mM FeCl3 + 2.5 mM K3Fe(CN)6 + 0.1 M KCl + 0.1 M HCl and applying a constant potential of +0.4 V for 4 min. Afterwards, the electrode was placed into a supporting electrolyte solution made of 0.1 M KCl + 0.1 M HCl and electrochemically activated by cycling between a potential range of 0.05 and +0.4 V at a scan rate of 50 mV/s for 25 cycles. To immobilize the enzyme GOx, a solution of the mixture of 20 lL of 9 mg mL1 of GOx and 10 lL of 2 mg mL1 of chitosan was cast onto the PB modified nanoporous Au electrode (Ti/ NPAu/PB/GOx). Chitosan was used to enhance the stability of the biosensor [21,22]. For comparison, a glucose biosensor based on a PB modified Au thin film (Ti/Au/PB/GOx) was also fabricated with the identical procedure. The Au thin film was coated onto the Ti substrate by argon plasma sputtering for 30 s. All prepared enzyme electrodes were stored at 4 °C when not in use.

200 150 100

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Fig. 1. The typical SEM image of the nanoporous Au coating (a) and the corresponding EDS spectrum (b).

surface area than the sputtered Au thin film. The oxidation and reduction peaks originate from the redox process of Prussian white (Fe43+[Fe2+(CN)6]3) and Prussian blue (K4Fe43+[Fe2+(CN)6]3) [23,24]. After we immobilized the enzyme GOx on the Ti/NPAu/PB surface, the redox peaks decreased as seen in Plot c. The immobilized GOx blocks some of the redox reaction sites, indicating the successful immobilization of GOx on the PB modified nanoporous Au surfaces. 3.3. Response of the glucose biosensor The electrochemical characterization of the biosensor was investigated by cyclic voltammetry. Fig. 2B presents the CV curves recorded in 0.1 M PBS (pH = 7.0) at a scan rate of 20 mV/s in the absence of glucose (Dashed line) and in the presence of 1 mM glucose (Solid line). It is seen that the reduction current increases while the oxidation current decreases, which relates to the oxidation of glucose by GOx catalysis. The mechanism for the reaction for the detection of glucose can be defined as follows [23]:

D-glucose þ O2 ¢ gluconolactone þ H2 O2  2þ 3þ 2þ K4 Fe2þ 4 ½Fe ðCNÞ6 3 þ 2H2 O2 ¢ Fe ½Fe ðCNÞ6 3 þ 4OH 2þ þ 2þ 2þ  Fe3þ 4 ½Fe ðCNÞ6 3 þ 4e þ 4K ¢ K4 Fe4 ½Fe ðCNÞ6 3

ð1Þ þ

þ 4K ð2Þ ð3Þ

The impact of the applied electrode potential and pH of the PBS on the amperometric response of the Ti/NPAu/PB/GOx electrode to glucose was investigated (data not shown), revealing the optimized sensing conditions: the applied potential of 0.1 V and pH = 7. Fig. 3a displays the steady-state amperometric response of (i) the Ti/Au/PB/GOx electrode and (ii) the Ti/NPAu/PB/GOx

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E/V (vs. Ag/AgCl) Fig. 2. Cyclic voltammograms of (A): the Ti/Au/PB electrode (Plot a), the Ti/NPAu/PB electrode (Plot b) and the Ti/NPAu/PB/Gox electrode (Plot c) recorded in a 0.1 M PBS (pH = 7.0) at a scan rate of 10 mV/s; (B) the Ti/NPAu/PB/GOx electrode in the absence (dashed line) and in the presence of 1 mM glucose (solid line).

biosensor recorded by successively adding glucose into the PBS under the optimized sensing conditions. For the Ti/NPAu/PB/GOx biosensor, 95% of the steady-state current was achieved at less than 15 s. The corresponding calibration curves for the two electrodes are shown in Fig. 3b. The amperometric response of the Ti/NPAu/ PB/GOx biosensor is much higher than that of the Ti/Au/PB/GOx electrode. A linear relationship of the current to the concentration of glucose was observed between 0 and 2.04 mM with a correlation coefficient of 0.9993 with 17 successive injections of 0.12 mM glucose. The sensitivity of the Ti/NPAu/PB/GOx biosensor is extremely high, 177 lA/mM. The detection limit is estimated to be 2.5 lM glucose (based on S/N = 3). To evaluate the biological activity of the immobilized GOx, the apparent Michaelis–Menten constant Km is generally used [25]. According to the Michaelis–Menten equation:

Imax =Is ¼ K m =C þ 1

ð4Þ

where I is the steady-state catalytic current, Imax is the maximum current measured under saturated conditions, C refers to the glucose concentration and Km stands for the apparent Michaelis–Menten constant. This equation allows us to plot the experimental 1/Is vs. 1/C to determine Km and Imax. According to our experimental data, Km was evaluated as 2.1 mM and the Imax was 443 lA. The va-

Fig. 3. (a) Amperometric response to successive additions of 0.12 mM glucose into a 0.1 M pH 7.0 PBS under the applied electrode potential 0.1 V; (b) the corresponding calibration curves for the response of: (i) the Ti/Au/PB/GOx and (ii) Ti/NPAu/PB/GOx electrode.

lue of Km is much lower than the reported 4.3 mM for gold nanoparticles [26], 19 mM for ZnO nanotubes [27], and 21.1 mM for microband gold electrodes [28]. The small Km value indicates that the immobilized GOx possesses high enzymatic activity and that the fabricated biosensor exhibits a high affinity for glucose.

3.4. Anti-interference and stability of biosensor We further studied the selectivity of the fabricated Ti/NPAu/PB/ GOx biosensor. The amperometric response of the biosensor to successive additions of common interfering substances and glucose in 0.1 M PBS (pH = 7.0) was measured under the applied electrode potential 0.1 V. As shown in Fig. 4, the biosensor shows no or very little response to ascorbic acid (AA), uric acid (UA), and acetaminophen (AP); in contrast, the biosensor exhibits very strong response to the successive injections of 0.1 mM glucose in the presence of 0.1 mM AA, 0.2 mM UA and 0.1 mM AP, indicating that the glucose biosensor fabricated in this study has a very high anti-interferent ability. The reproducibility of the biosensor was also examined by measuring the response towards 1 mM glucose in 50 ml pH 7.0 PBS. Four biosensors constructed independently gave a relative standard deviation (R.S.D.) of 4.2%. The long-term stability of the biosensor was evaluated by measuring its performance every week. The biosensor retained around 85% of the initial

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Acknowledgments Glucose

I/ µ A

0

Glucose AP

AA

UA

This work was supported by a Discovery Grant from the Natural Sciences and Engineering Research Council of Canada (NSERC). A. Chen acknowledges NSERC and the Canada Foundation of Innovation (CFI) for the Canada Research Chair Award in Material and Environmental Chemistry.

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t/s Fig. 4. Effect of interferents, 0.1 mM ascorbic acid (AA), 0.1 mM acetaminophen and 0.02 mM uric acid (UA), on the response of the Ti/NPAu/PB/GOx biosensor to successive additions of 0.1 mM glucose.

response after four weeks. After 60 days, the biosensor was still very active, retaining over 70% of the initial response. 4. Conclusions In this communication, we have demonstrated a novel glucose biosensor based on the highly efficient immobilization of GOx on a PB modified nanoporous Au surface. The nanoporous Au was directly grown through a facile hydrothermal method and modified by electrodeposited Prussian blue. Our study has shown that the PB modified nanoporous Au substrate provides an excellent matrix for the immobilization of GOx owing to its large surface area and strong adsorption ability for enzyme binding, high chemical stability, low inherent toxicity and high conductivity. The small Km value of the Ti/NPAu/PB/GOx biosensor indicates that the immobilized GOx possesses high enzymatic activity and that the fabricated biosensor exhibits a high affinity for glucose. Our electrochemical tests show that the fabricated Ti/NPAu/PB/GOx biosensor exhibits fast amperometric response, a low detection limit (2.5 lM), extremely high sensitivity (177 lA mM1), good reproducibility, high anti-interferent ability and long-time stability.

[1] S.R. Lee, K. Sawada, H. Takao, M. Ishida, Biosens. Bioelectron. 24 (2008) 650. [2] R. Antiochia, L. Gorton, Biosens. Bioelectron. 22 (2007) 2611. [3] T. Saxl, F. Khan, D. Matthews, Z.L. Zhi, S. Ameer-Beg, O. Rolinski, J. Pickup, Biosens. Bioelectron. 24 (2009) 3229. [4] H. Chung, M.A. Arnold, M. Rhiel, D.W. Murhammer, Appl. Spectrosc. 50 (1996) 270. [5] L. Yang, C. Du, X. Luo, J. Nanosci. Nanotechnol. 9 (2009) 2660. [6] A.C. Mendes, M.M. Caldeira, C. Silva, S.C. Burgess, M.E. Merritt, F. Gomes, C. Barosa, T.C. Delgado, F. Franco, P. Monteiro, L. Providencia, J.G. Jones, Magn. Reson. Med. 56 (2006) 1121. [7] L.A. Hammad, M.M. Saleh, M.V. Novotny, Y. Mechref, J. Am. Soc. Mass Spectrom. 20 (2009) 1224. [8] A.P. Cook, T.M. MacLeod, J.D. Appletont, A.F. Fell, J. Clin. Pharm. Ther. 14 (2009) 189. [9] I. Rzygalinski, E. Pobozy, R. Drewnowska, M. Trojanowicz, Electroanalysis 29 (2008) 1741. [10] J. Lewandowski, P.S. Malchesky, J. Krzymien, E. Szczepanskasadowska, M. Nalecz, Y. Nose, IEEE Trans. Biomed. Eng. 31 (1984) 582. [11] P. Benvenuto, A.K.M. Kafi, A. Chen, J. Electroanal. Chem. 627 (2009) 76. [12] Y.G. Zhou, S. Yang, Q.Y. Qian, X.H. Xia, Electrochem. Commun. 11 (2009) 216. [13] J. Wang, D.F. Thomas, A. Chen, Anal. Chem. 80 (2008) 997. [14] R.T. Kachoosangi, G.G. Wildgoose, R.G. Compton, Analyst 133 (2008) 888. [15] P. Holt-Hindle, S. Nigro, M. Asmussen, A. Chen, Electrochem. Commun. 10 (2008) 1438. [16] A. Umar, M.M. Rahman, A. Al-Hajrya, Y.B. Hahn, Electrochem. Commun. 11 (2009) 278. [17] A.K.M. Kafi, G. Wu, A. Chen, Biosens. Bioelectron. 24 (2008) 566. [18] Y. Ofir, B. Samanta, V.M. Rotello, Chem. Soc. Rev. 37 (2008) 1814. [19] X. Wang, H. Gu, F. Yin, Y. Tu, Biosens. Bioelectron. 24 (2009) 1527. [20] Y. Liu, Z.Y. Chu, W.Q. Jin, Electrochem. Commun. 11 (2009) 484. [21] R. Khan, M. Dhayal, Electrochem. Commun. 10 (2008) 263. [22] A.K.M. Kafi, A. Chen, Talanta 79 (2009) 97. [23] F. Ricci, G. Palleschi, Biosens. Bioelectron. 21 (2005) 389. [24] A. Chen, E.I. Rogers, R.G. Compton, Electroanalysis 21 (2009) 29. [25] F.R. Shu, G.S. Wilson, Anal. Chem. 48 (1976) 1679. [26] S. Zhang, N. Wang, H. Yu, Y. Niu, C. Sun, Bioelectrochemistry 67 (2005) 15. [27] T. Kong, Y. Chen, Y. Ye, K. Zhang, Z. Wang, X. Wang, Sens. Actuators B 138 (2009) 344. [28] H. Ju, D. Zhou, Y. Xiao, H. Chen, Electroanalysis 10 (1998) 541.