Glucose biosensor based on immobilization of glucose oxidase in poly(o-aminophenol) film on polypyrrole-Pt nanocomposite modified glassy carbon electrode

Biosensors and Bioelectronics 22 (2007) 2898–2905

Glucose biosensor based on immobilization of glucose oxidase in poly(o-aminophenol) film on polypyrrole-Pt nanocomposite modified glassy carbon electrode Jing Li, Xiangqin Lin ∗ Department of Chemistry, University of Science and Technology of China, # 96 Jinzhai Rd., Hefei, Anhui 230026, PR China Received 22 August 2006; received in revised form 30 November 2006; accepted 1 December 2006 Available online 9 January 2007

Abstract Novel Pt nanoclusters embedded polypyrrole nanowires (PPy-Pt) composite was electrosynthesized on a glassy carbon electrode, denoted as PPyPt/GCE. A glucose biosensor was further fabricated based on immobilization of glucose oxidase (GOD) in an electropolymerized non-conducting poly(o-aminophenol) (POAP) film that was deposited on the PPy-Pt/GCE. The morphologies of the PPy nanowires and PPy-Pt nanocomposite were characterized by field emission scanning electron microscope (FE-SEM). Effect of experimental conditions involving the cycle numbers for POAP deposition and Pt nanoclusters deposition, applied potential used in glucose determination, temperature and pH value of the detection solution were investigated for optimization. The biosensor exhibited an excellent current response to glucose over a wide linear range from 1.5 × 10−6 to 1.3 × 10−2 M (r = 0.9982) with a detection limit of 4.5 × 10−7 M (s/n = 3). Based on the combination of permselectivity of the POAP and the PPy films, the sensor had good anti-interference ability to ascorbic acid (AA), uric acid (UA) and acetaminophen. The apparent Michaelis–Menten constant (Km ) and the maximum current density (Im ) were estimated to be 23.9 mM and 378 ␮A/cm2 , respectively. In addition, the biosensor had also good sensitivity, stability and reproducibility. © 2006 Elsevier B.V. All rights reserved. Keywords: Polypyrrole nanowires; Platinum nanoclusters; Composite; Poly(o-aminophenol); Glucose biosensor; Electrochemical deposition

1. Introduction Recently, the development of glucose biosensor has been received considerable attention because determination of glucose concentration is very important in clinical applications (Yoshimura and Hozumi, 1996; Mizutani and Yabuki, 1997; Rigby et al., 1999; Wilson and Hu, 2000). Most glucose measurements are based on immobilization of glucose oxidase (GOD) for detecting H2 O2 concentration which is produced from the GOD enzymatic reaction. Since GOD can recognize glucose target molecules quickly and accurately in complicated systems, a proper matrix on the electrode surface should be well designed for immobilization of GOD while maintaining its highly enzymatic activity. In recent years, the use of nano-materials has been widely interested due to their catalytic ability and good biocom-

∗

Corresponding author. Tel.: +86 551 3606646; fax: +86 551 3601592. E-mail address: [email protected] (X. Lin).

0956-5663/$ – see front matter © 2006 Elsevier B.V. All rights reserved. doi:10.1016/j.bios.2006.12.004

patibility (Liu and Chuang, 2003; Majid et al., 2006; Ren et al., 2005; Sulak et al., 2006). Extensive efforts have been devoted to improve the sensitivity, long-term stability and anti-interference ability of the biosensor. The entrapment of GOD in electropolymerized film is a simple and efficient way to develop glucose sensors with high sensitivity, good stability and fast responsibility. The entrapping technique should control the thickness of the layer and the loading of the enzyme, while keeping the enzyme highly bioactive. The permselective polymer films used are basically non-conductive, in order to reject electroactive species such as ascorbic acid (AA) and uric acid (UA), which are coexisting in blood samples. GOD entrapped in poly(m-phenylenediamine) multi-layer films at the platinum electrode has been reported. The apparent Michaelis–Menten constant (Km ) and the maximum steady-state current density (Im ) of the sensor were reported as 24 mM and 80 ␮A/cm2 , respectively (Yang et al., 2002). A glucose sensor for in vivo monitoring was developed based on immobilization of GOD at a platinized diamond

J. Li, X. Lin / Biosensors and Bioelectronics 22 (2007) 2898–2905

microfiber electrode. The sensor had a linear range of 1–70 mM glucose for the determination (Olivia et al., 2004). The immobilization of GOD in poly(3,4-ethylenedioxythiophene) (PEDT) films has also been reported (Piro et al., 2001). Poly(oaminophenol) (POAP) film has been successfully used to develop UA sensor (Miland et al., 1996; Pan et al., 2006), cholesterol (Vidal et al., 2003), catechol (Mu, 2006) and glucose sensor (Nakabayashi et al., 1998; Zhang et al., 1996), providing an efficient barrier to exclude interferents and to protect the electrode from fouling. The OAP film can be in situ electropolymerized on the electrode surface, and the thickness can be controlled within 10–100 nm due to the self-limiting effect, while providing good stability and anti-interference ability. However, the non-conducting polymer has also a matter of concern about their relatively high detection limit and low current responsibility (Malitesta et al., 1990), so conductive polymers have been reviewed as promising support for fabrication of biosensors. Polypyrrole (PPy), a key member of organic conducting polymers, can serve as the enzyme-hosting matrix for bio-molecules (Cosnier et al., 1999; Ramanavicius et al., 1999; Gaspar et al., 2001; Vidal et al., 2003; Retama et al., 2004; Adeloju et al., 2005), owing to its easy preparation, high conductivity and good stability. A glucose sensor based on immobilization of GOD in PPy and Prussian Blue bilayers has been reported with high sensitivity and fast response (Garjonyte and Malinauskas, 2000). Nevertheless, the morphology, spatial structure and specific surface area have been recognized as the most important factors influencing the catalytic performance of the PPy film. Due to its well ordered polymer chain structure with high aspect ratio and small cross dimensions, PPy nanowire has been received significant attention, which can be generated on the electrode surface under stationary potentials by template-free method (Tian et al., 2005). Further deposition of platinum by cyclic voltammetry (CV) method can generate Pt nanoclusters embed in the PPy nanowires, forming PPy-Pt composite matrix. This polymer-metal nanocomposite matrix can provide a highly porous structure with large effective surface area, good electronic conductivity and high catalytic activity, which are beneficial to immobilization of enzymes for fabrication of highly sensitive biosensors. This paper described the fabrication and characterization of a glucose sensor based on the PPy-Pt nanocomposite modified glassy carbon electrode, which was covered by a layer of POAP-GOD film. The sensor exhibited excellent performances, such as low detection limit, wide linear range, large current density, quick current response, high sensitivity, good stability and reproducibility. Particularly, the interferences of AA, UA and acetaminophen can be completely eliminated, due to the combination of permselective films of POAP and PPy. 2. Experimental 2.1. Reagents Pyrrole was obtained from Aldrich and purified twice by distillation under the protection of high purity nitrogen and then

2899

kept in a refrigerator before use. Ascorbic acid (AA), uric acid (UA), acetaminophen, K2 PtCl6 , and o-aminophenol (OAP) were obtained from Chemical Reagent Company of Shanghai (Shanghai, China). OAP was purified by recrystallization before use. Glucose oxidase (GOD, 300 U/mg) and d-glucose were obtained from Sigma. All other Chemicals were of analytical-reagent grade and used without further purification. Doubly distilled water and high purity N2 were used. The 0.1 M phosphate buffer solutions of different pH values (PBSs) were prepared for the study. 2.2. Apparatus and sensing protocols All electrochemical measurements were performed with electrochemistry workstation CHI 660A (ChenHua Instruments Co., Shanghai, China). A conventional three-electrode system was used, which was composed of a working electrode, a platinum wire counter electrode, and a saturated calomel reference electrode (SCE). A glassy carbon disk working electrode (GCE, Φ4.0 mm) was used as the basal electrode for fabrication. FE-SEM images of the electrode surfaces were obtained on JSM-6700F field emission scanning electron microanalyser (FE-SEM) (JEOL, Japan). X-ray photoelectron spectroscopy (XPS) measurements were performed on an ESCALAB MK2 spectrometer (VG Co., UK) with Mg K-alpha radiation as the source for excitation. Powder X-ray diffraction (XRD) spectra were recorded on a MXPAHF rotating anode Xray diffractometer (Japan) with Cu K␣ radiation source ˚ (λ = 1.54056 A). The glucose response of enzyme electrodes were measured based on the electrooxidation current of H2 O2 that was produced during the enzymatic reaction. Amperometric response of the enzyme electrodes were determined in a one-compartment glass cell with a magnetic stirrer containing 25 ml air-saturated 0.1 M PBS. After the background current reached a steadystate value at an applied potential, glucose was injected into the solution using a micro-syringe and the current was recorded. Unless otherwise stated, all experiments were carried out at 25 ◦ C. 2.3. Synthesis of PPy nanowires Prior to modification, the basal GCE was polished to a mirror finish using alumina slurries with different powder size down to 0.05 ␮m. After each polishing, the electrode was sonicated in ethanol and doubly distilled water for 5 min, successively, in order to remove any adsorbed substances on the electrode surface. Finally, it was dried under nitrogen atmosphere ready for use. The PPy was electrochemically deposited at a constant potential of 0.80 V for 120 s in an aqueous solution of 0.1 M LiClO4 and 0.1 M carbonate containing 0.15 M pyrrole. The electrode was then transferred into 0.1 M HClO4 solution for 12 h aging to remove any carbonate ions (Tian et al., 2005). The PPy nanowires modified electrode was obtained and denoted as PPy/GCE.

2900

J. Li, X. Lin / Biosensors and Bioelectronics 22 (2007) 2898–2905

2.4. Synthesis of Pt nanoclusters Platinum nanoclusters were electrochemically deposited by cyclic voltammetry (CVs) in 0.5 M H2 SO4 solution containing 2 mM K2 PtCl6 with a potential scanning from 0.40 to −0.25 V at 50 mV/s for 30 cycles. The PPy-Pt nanocomposite modified electrode was obtained and denoted as PPy-Pt/GCE. For a comparison, the same Pt deposition process was conducted at bare GCE and denoted as Pt/GCE. 2.5. Fabrication of glucose sensor After synthesizing the PPy nanowires and Pt nanoclusters on the electrode, the designed glucose sensor was fabricated by electrodeposition of OAP on the PPy-Pt/GCE in 0.1 M pH 5.6 acetate buffer containing 5.0 mM OAP and 2.0 mg/ml GOD by potential scanning between −0.20 to 0.80 V at 50 mV/s for 15 cycles. The immobilization of GOD in the polymer film occurred together with the electropolymerization of OAP. The prepared electrode was denoted as POAP-GOD/PPy-Pt/GCE. For a comparison, the same POAP-GOD deposition was carried out at Pt/GCE and PPy/GCE, and the obtained electrodes were denoted as POAP-GOD/Pt/GCE and POAP-GOD/PPy/GCE, respectively. All the enzyme electrodes were stored in 0.1 M PBS (pH 7.0) at 4 ◦ C before use.

a cucumber-like surface, the size of which could not be seen clearly since the PPy cover is not transparent. The growing of large embedded Pt nanoclusters generated “a string of beads” like structure, which can be clearly seen on the top of the matrix. XPS spectra were obtained for PPy-Pt/GCE. The spectra shows two bands appear at 70.8 and 74.2 eV, corresponding to the Pt 4f7/2 and Pt 4f5/2 signals, respectively. This is consistent with the Pt(0) formation. The powder XRD determination was used to characterize the crystal structure of Pt nanoparticles, as shown in Fig. S1. Four peaks appears at 2θ of 39.8◦ , 46.2◦ , 67.5◦ and 82.2◦ in the range of 20–90◦ , which can be assigned to the (1 1 1), (2 0 0), (2 2 0), (3 1 1) structure of the face-centered cubic lattice of Pt(0), respectively. The particle size of Pt is calculated according to the Scherrer formula (Radmilovic et al., 1995), L = 0.9 ␭/B(2θ) cos θ B , where L is the average size of particles, ␭, the X-ray wavelength, B(2θ) , full-width half-maximum of diffraction peak, and θ B is the angle of the peak maximum. Based on the reflection peak at 2θ of 39.8◦ , an averaged diameter of the Pt particles is calculated as 9.1 nm. The standard deviation is about 2.5%. The results indicate the Pt nanoparticles of about 100 nm size embedded in the PPy nanowires are clusters of nanocrystallites of about 9.1 nm in diameter. This phenomenon is quite different from the report that the metal nanoparticles tend to congregate into large nanostructures at the surface of PPy film (Li and Shi, 2005).

3. Results and discussion 3.2. Electrochemical deposition of POAP-GOD film 3.1. Characterization of electrode surface The morphologies of the PPy nanowires and PPy-Pt nanocomposite were characterized by FE-SEM, as shown in Fig. 1. Fig. 1A shows the sponge-like matrix of PPy nanowires on GCE. The wire diameter is about 60 nm. The three-dimensional (3D) structure of the matrix has large number of gaps and pores, which are beneficial to Pt deposition and enzyme incorporation. Fig. 1B shows the morphology of the PPy-Pt nanocomposite, which is generated by the CV deposition of Pt on the PPy/GCE. It can be seen that the deposited Pt nanoclusters are not on the surface of PPy but embed in PPy, generating the PPy-Pt nanocomposite. The PPy-Pt composite has also a matrix structure similar to the PPy nanowires matrix, however, the diameter of the PPy nanowire expands to 100–120 nm due to Pt embedding. The small Pt nanoclusters embedded PPy nanowires form

Electrochemical polymerization of OAP has been studied extensively in recent years (Jackowska et al., 1993; Lobo et al., 1996; Tao et al., 2004). It is well known that the POAP film formed in acidic media would be electroconductive, but nonconductive in neutral and basic solutions (Barbero et al., 1989). To avoid denaturation of the enzyme, the deposition of POAPGOD was performed in pH 5.6 acetate buffer solution to obtain non-conductive polymer film for the biosensor construction. Fig. 2 presents the multi-cycle CVs of POAP-GOD deposition at GCE, PPy/GCE, Pt/GCE and PPy-Pt/GCE for comparison. During OAP electropolymerization, GOD molecules are desirable to be readily incorporated into the polymer film from the solution. The oxidation wave of OAP at bare GCE is shown in Fig. 2A. A broad irreversible anodic wave appears in the potential range from +0.17 to +0.80 V at the first

Fig. 1. FE-SEM images of the PPy nanowires (A) and the PPy-Pt nanocomposite (B) on bare GCE.

J. Li, X. Lin / Biosensors and Bioelectronics 22 (2007) 2898–2905

2901

Fig. 2. Multi-cycle CVs of electropolymerization of POAP-GOD on bare GCE (A), PPy/GCE (B), Pt/GCE (C), and PPy-Pt/GCE (D) at 50 mV/s in 0.1 M acetate buffer (pH 5.6) + 5.0 mM OAP + 2.0 mg/mL GOD.

cycle and is decreasing quickly at the following cycles. The peak potential shifts to positive direction in the continuous cycling and the peak current drops significantly with each scan until it reaches a minimum value, which is similar to the case for simple OAP electropolymerization, indicating a compact and insulating film is formed on the electrode surface. Similar phenomena are observed at the PPy/GCE (Fig. 2B). A visible platinum film formed on the bare electrode surface after Pt deposition, which has a superior catalytic activity toward OAP oxidation. As shown in Fig. 2C, the Pt/GCE gives well formed oxidation peak of OAP at about 0.27 V in the first CV cycle. Although the PPy-Pt/GCE also presents a well oxidation peak of OAP at 0.35 V at the first CV cycle (Fig. 2D), the peak potential is 80 mV more positive than that at Pt/GCE. The mass of deposited POAP can be calculated through the integrated charge (Q) passed during electropolymerization based on Faradays law. According to the equation of Q = nFAΓ (Laviron, 1979), considering a value of 2 for n and 109 for OAP molecular weight. Assuming a value of about 1.30 g/cm3 for POAP density at 25 ◦ C, the POAP film thickness can easily be calculated. It showed that the POAP film thickness was 8 nm on bare GCE, 10 nm on PPy/GCE, 20 nm on Pt/GCE and 25 nm on PPy-Pt/GCE for 15 cycles deposition. The amount of GOD entrapped in the film could be assumed as proportional to the formal thickness of the POAP film. The maximum current response at the POAP-GOD/PPy-Pt/GCE was obtained when the polymerization cycle number for POAP-GOD was 15. The current response to AA, UA and acetaminophen decreased quickly with increase of the cycle number from 5 and then reached a minimum value. However, the response time was increased with

further increasing the cycle number. Thus, the optimal cycle number of 15 was used to fabricate the enzyme electrodes. 3.3. Effect of the applied potential The effect of the applied potential on the current response of the POAP-GOD/PPy-Pt/GCE was examined. As shown in Fig. S2A, the current response starts to increase at potential of 0.35 V and reaches a maximum value at 0.60 V. In order to further demonstrate the mechanism of the detected current, the PPy-Pt/GCE was used to determine H2 O2 . The CV of 2.0 mM H2 O2 at the PPy-Pt/GCE is obtained and shows in Fig. S3. A broad anodic peak appears at about 0.5 V at scan rate of 50 mV/s. The oxidation current of H2 O2 increases from the potential range of 0.3–0.5 V, demonstrating that the PPy-Pt/GCE can electrocatalytically oxidize H2 O2 . As reported before, the determination of glucose based on the enzyme electrodes always set a potential at 0.60 V (Zhou et al., 2005; Ye et al., 2005; Pan et al., 2004; Trojanowicz and Miernik, 2001), in order to obtain relatively large response current and shorten the response time. Thus, we choose the operational potential of 0.60 V in our experiments. 3.4. Effect of the pH The pH value is another variable that affects the current response of the enzyme electrodes. The pH dependence of the response of the POAP-GOD/PPy-Pt/GCE has been investigated using 0.1 M PBS and the corresponding result showed in Fig. S2(B). The current increased from pH 4.5–7.0, while decreased sharply above pH 7.0. The maximum current of the

2902

J. Li, X. Lin / Biosensors and Bioelectronics 22 (2007) 2898–2905

enzyme electrode was obtained at pH 7.0, due to the entrapment of GOD in the POAP film covered on the PPy-Pt nanocomposite, which made GOD more active in neutral solution. We also investigated free enzyme toward glucose oxidation using the PPy-Pt/GCE under the same solution conditions. The maximum current value was also obtained at pH 7.0, which indicated that the deposition of POAP-GOD do not change the bioactivity of GOD. Considering the practically application of GOD, neutral buffer solution (pH 7.0) was selected for glucose detection. 3.5. Effect of the temperature

Scheme 1. Proposed mechanism of the POAP-GOD/PPy-Pt/GCE polarized at 0.60 V vs. SCE.

The temperature is also an important factor for the activity of enzyme. The response of the enzyme electrode was measured at temperature between 10 and 60 ◦ C. As shown in Fig. S2(C), the current response increases almost linearly with temperature from 10 to 40 ◦ C, but decreases linearly from 40 to 60 ◦ C. The later is due to the deactivation of GOD. However, the former is attributed to the increase of the activity of the immobilized GOD. Thus, the activation energy of the immobilized GOD can be calculated as 25.9 kJ/mol based on the Arrhenius formula. According to the results, we can conclude that the PPy-Pt nanocomposite and POAP film can offer a good environment for GOD, which make the sensor more stable at high temperature. Considering the convenience of the practical application, 25 ◦ C was selected in our experiment. 3.6. Amperometric determination of glucose The designed sensor POAP-GOD/PPy-Pt/GCE has sensitive response to glucose. Fig. S4(A) shows a typical current-time curve of the POAP-GOD/PPy-Pt/GCE at 0.60 V to the successive addition of glucose in a stirred solution. The response current increases with increasing the concentration of glucose and finally reaches to a steady-state value. A fast response time of about 7 s was estimated. Such a fast response can be attributed to the thin POAP film and the high dispersion of Pt nanoclusters embedded. The typical amperometric response curve of the sensor was plotted in Fig. S4(B) (curve a), together with that for POAP-GOD/Pt/GCE (curve b) and POAP-GOD/PPy/GCE (curve c) for comparison. The response current of the POAP-GOD/PPy-Pt/GCE (curve a) was linear with glucose concentration to 13 mM with a detection limit of 4.5 × 10−7 M (s/n = 3), for concentrations higher than 13 mM, the response gradually reached to a plateau. The linear ranges up to a value 12 mM of glucose concentration were also obtained for

POAP-GOD/Pt/GCE and POAP-GOD/PPy/GCE, respectively. The detection limit, linear range, current sensitivity, reproducibility (RSD) and maximum current density for 10 repeated determinations of these sensors were summarized in Table 1. The POAP-GOD/PPy-Pt/GCE has lower detection limit and higher sensitivity, owing to the application of PPy-Pt nanocomposite. Due to the 3D conducting nanowire matrix of PPy, the effective surface area of the electrode was significantly enlarged, so the surface coverage of GOD on the electrode can be also remarkably enhanced. On the other hand, the embedded Pt nanoclusters that consist of 9.1 nm sized nanocrystallites can provide even higher active surface area for catalytic oxidation of H2 O2 . Furthermore, the nano-thick PPy coverage of Pt nanoclusters can provide close contact between the embedded GOD and the Pt catalytic site, thus enhance the catalytic efficiency and shorten the response time. The PPy coverage could also stabilize the Pt nanoclusters, leading to stabilization of the sensor activity. Scheme 1 shows the proposed mechanism for the POAPGOD/PPy-Pt/GCE polarized at 0.60 V versus SCE. The added glucose transferred fast into the POAP film so as to be oxidized by the entrapped GOD in the presence of O2 , according to the following reaction: GOD

glucose + O2 −→gluconic acid + H2 O2

(1)

The generated H2 O2 can penetrate through the POAP film and be electrocatalytically oxidized at the PPy-Pt conducting nanocomposite, in result of anodic current according to: PPy-Pt

H2 O2 −→ O2 + 2H+ + 2e−

(2)

when the enzyme electrode is immersed in glucose solutions, the reactions (Eqs. (1) and (2)) should occur. The sensitive detection of H2 O2 is essential for the sensitivity of the glucose sensor.

Table 1 Analytical characteristics of the glucose sensorsa Sensorb

Detection limit (s/n = 3) (␮M)

Linear range (mM)

Sensitivity (mA/M cm2 )

R.S.D.c (N = 10) (%)

Km (mM)

Maximum current density (␮A/cm2 )

#1 #2 #3

0.45 0.90 0.95

0.0015–13 (r = 0.9982) 0.0042–12 (r = 0.9981) 0.0055–12 (r = 0.9980)

9.9 5.5 3.5

2.8 3.4 3.2

23.9 10.2 6.9

378 156 122

a b c

Determined in 0.1 M pH 7.0 PBS at 0.60 V vs. SCE. Sensors #1 is the POAP-GOD/PPy-Pt/GCE; #2, the POAP-GOD/Pt/GCE; #3, the POAP-GOD/PPy/GCE. Determined for 1.0 mM glucose.

J. Li, X. Lin / Biosensors and Bioelectronics 22 (2007) 2898–2905

According to the Lineweaver–Burk form of the Michaelis–Menten equation (Shu and Wilson, 1976), the relation between the reciprocal of the response current (i−1 s ) and the reciprocal of glucose concentration (Cg−1 ) was obtained. The Michaelis–Menten constant (Km ) and the maximum current density (Im ) of the POAP-GOD/PPy-Pt/GCE were calculated to be 23.9 mM and 378 ␮A/cm2 , respectively. These values for the POAP-GOD/Pt/GCE and POAP-GOD/PPy/GCE were also calculated and equal to 10.2 mM and 156 ␮A/cm2 , 6.9 mM and 122 ␮A/cm2 , correspondingly. These parameters also listed in Table 1 for comparison. Generally, the Km was used to evaluate enzyme activity. In our experiment system, the Km value should be also determined by the high enzyme loading and PPy-Pt nanocomposite electrocatalytic activity, in result of the Km of the POAP-GOD/PPy-Pt/GCE is higher than the other enzyme electrodes. Since the higher Km is, the larger Im will be, the Km value can be considered as a criterion for the effective enzyme loading and the electrocatalytic activity for H2 O2 oxidation. 3.7. Interference test Fig. 3 shows CVs obtained in the absence (a) and presence of AA (b), UA (c) and acetaminophen (d) at bare GCE, PPy/GCE, POAP-GOD/Pt/GCE, and POAP-GOD/PPy-Pt/GCE, respectively. It was found that the bare GCE responds well to all these electroactive species (Fig. 3A), so it could not be used to detect glucose. After the formation of PPy nanowires on GCE surface, the oxidation current of acetaminophen decreases significantly (Fig. 3B, curve d). The irreversible oxidation peak of

2903

UA (Fig. 3B, curve c) appears at 0.4 V and the current decreases to about 40%. Furthermore, the response current of AA (Fig. 3B, curve b) do not diminish. These results indicate that the oxidation of AA and UA is considerable at the PPy/GCE. After POAP-GOD film coating on the Pt/GCE, the oxidation current response for all these species (especially AA and UA) is significantly depressed (Fig. 3C, curve b and c), however, the response current of acetaminophen (Fig. 3C, curve d) increases at potentials more positive than 0.4 V in the anodic scan. Their interferences were still beyond ignoring. Fig. 3D showed the CVs of these interferents at the POAP-GOD/PPy-Pt/GCE. It can be clearly seen that the current response to all these species have been completely eliminated. Consequently, it demonstrates that though both the electropolymerized films are permeable to these interferents, the permeability to these interferents can be significantly lowered by the combination of the POAP and PPy films. Fig. 4 shows amperometric response of POAP-GOD/PPyPt/GCE (A) and POAP-GOD/Pt/GCE (B) to the consecutive addition of glucose (a), UA (b), AA (c), acetaminophen (d) and glucose (e and f) to the solution. The current response of 5.5 mM glucose (a) can be observed at POAP-GOD/PPyPt/GCE (Fig. 4A) and POAP-GOD/Pt/GCE (Fig. 4B), whereas the later response is much smaller than the former. For successively adding of 0.5 mM UA (b), 0.2 mM AA (c) and 2.0 mM acetaminophen (d) in the glucose solution, the POAP-GOD/PPyPt/GCE has nearly no current response to these additions (Fig. 4A). The POAP-GOD/Pt/GCE can also eliminate the influence of UA and AA, however, 2.0 mM acetaminophen has

Fig. 3. CVs at bare GCE (A), PPy/GCE (B), POAP-GOD/Pt/GCE (C) and POAP-GOD/PPy-Pt/GCE (D) at 50 mV/s in 0.1 M PBS (a), a + 0.2 mM AA (b), a + 1.0 mM UA (c), and a + 2.0 mM acetaminophen (d).

2904

J. Li, X. Lin / Biosensors and Bioelectronics 22 (2007) 2898–2905

Fig. 4. Amperometric response to the injection of 5.5 mM glucose (a), 0.5 mM UA (b), 0.2 mM AA (c), 2.0 mM acetaminophen (d), 0.9 mM glucose (e), and 1.5 mM glucose (f) at POAP-GOD/PPy-Pt/GCE (A) and POAP-GOD/Pt/GCE (B) in 0.1 M pH 7.0 PBS.

significant response (Fig. 4B). The interference of these electroactive compounds to the glucose response was also examined by finally adding glucose into the mixed solution. The influence of UA, AA and acetaminophen to the glucose response was little at the POAP-GOD/PPy-Pt/GCE in comparison with the POAP-GOD/Pt/GCE. The amperometric data coincide with the phenomena observed in the CV study (Fig. 3). These facts indicate that the bilayer structure of PPy and POAP films is useful to avoid interference and the POAP-GOD/PPy-Pt/GCE has superior anti-interference ability.

Acknowledgements

3.8. Stability of the POAP-GOD/PPy-Pt/GCE

References

Stability is a basic requirement for fabrication of glucose sensors. Generally, denaturation and loss of enzyme could occur during the storage of the sensor, in result of decreasing the sensitivity. The storage stability of the designed sensor POAPGOD/PPy-Pt/GCE was investigated. For detection of 5.0 mM glucose, there was no significantly decrease in current response in the first 7 days by every day use and then stored in 0.1 M PBS (pH 7.0) at 4 ◦ C. Only about 11% decrease occurred after 30 days. A 76% response current was still retained after 60 days. The results imply that the GOD molecules entrapped in the POAP film are stable and can retain bioactivity well. This should be attributed to the PPy-Pt nanocomposite structure, which provided biologically compatible matrix for immobilizing the POAP-GOD film.

Adeloju, S.B., Ohanessian, A., Duc, N.N., 2005. Synth. Met. 153 (1–3), 17–20. Barbero, C., Silber, J.J., Sereno, L., 1989. J. Electroanal. Chem. 263 (2), 333–352. Cosnier, S., Senillou, A., Gratzel, M., Comte, P., Vlachopoulos, N., Renault, N.J., Martelet, C., 1999. J. Electroanal. Chem. 469 (2), 176–181. Garjonyte, R., Malinauskas, A., 2000. Sens. Actuator B 63 (1–2), 122–128. Gaspar, S., Habermuller, K., Csoregi, E., Schuhmann, W., 2001. Sens. Actuator B 72 (1), 63–68. Jackowska, K., Bukowska, J., Kudelski, A., 1993. J. Electroanal. Chem. 350 (1–2), 177–187. Laviron, E., 1979. J. Electroanal. Chem. 100 (1–2), 263–270. Li, Y., Shi, G.Q., 2005. J. Phys. Chem. B 109 (50), 23787–23793. Liu, Y.C., Chuang, T.C., 2003. J. Phys. Chem. B 107 (45), 12383–12386. Lobo, M.J., Miranda, A.J., LopezFonseca, J.M., Tunon, P., 1996. Anal. Chim. Acta 325 (1–2), 33–42. Majid, E., Hrapovic, S., Liu, Y.L., Male, K.B., Luong, J.H.T., 2006. Anal. Chem. 78 (3), 762–769. Malitesta, C., Palmisano, F., Torsi, L., Zambonin, P.G., 1990. Anal. Chem. 62 (24), 2735–2740. Miland, E., Ordieres, A.J.M., Blanco, P.T., Smyth, M.R., Fagain, C.O., 1996. Talanta 43 (5), 785–796. Mizutani, F., Yabuki, S., 1997. Biosens. Bioelectron. 12 (9–10), 1013–1020. Mu, S.L., 2006. Biosens. Bioelectron. 21 (7), 1237–1243. Nakabayashi, Y., Wakuda, M., Imai, H., 1998. Anal. Sci. 14 (6), 1069–1076. Olivia, H., Sarada, B.V., Honda, K., Fujishima, A., 2004. Electrochim. Acta 49 (13), 2069–2076. Pan, D.W., Chen, J.H., Nie, L.H., Tao, W.Y., Yao, S.Z., 2004. Electrochim. Acta 49 (5), 795–801. Pan, X.H., Zhou, S., Chen, C., Kan, J.Q., 2006. Sens. Actuator B 113 (1), 329–334. Piro, B., Dang, L.A., Pham, M.C., Fabiano, S., Tran-Minh, C., 2001. J. Electroanal. Chem. 512 (1–2), 101–109. Radmilovic, V., Gasteiger, H.A., Ross, P.N., 1995. J. Catal. 154 (1), 98–106. Ramanavicius, A., Habermuller, K., Csoregi, E., Laurinavicius, V., Schuhmann, W., 1999. Anal. Chem. 71 (16), 3581–3586.

4. Conclusions In this work, a novel glucose sensor was fabricated by electrochemical deposition of POAP-GOD film on PPy-Pt nanocomposite modified GCE. In comparison with POAP-GOD film deposited on Pt/GCE and PPy/GCE, the designed sensor POAP-GOD/PPy-Pt/GCE has higher current sensitivity with a wider linear range, lower detection limit, larger apparent Km and Im , good anti-interference ability and excellent stability. The advantage of the designed sensor can be attributed to the application of the PPy-Pt nanocomposite matrix, which provides a special porous, biocompatible and highly catalytic activity. The nanostructure of Pt nanoclusters embedded in PPy nanowires is favorable for fabrication of biosensors.

The authors appreciate the financial support from National Natural Science Foundation of China (no. 20575062) and the Specialized Research Fund for the Doctoral Program of Higher Education (no. 20040358021). Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at doi:10.1016/j.bios.2006.12.004.

J. Li, X. Lin / Biosensors and Bioelectronics 22 (2007) 2898–2905 Ren, X.L., Meng, X.W., Tang, F.Q., 2005. Sens. Actuator B 110 (2), 358–363. Retama, J.R., Cabarcos, E.L., Mecerreyes, D., Lopez-Ruiz, B., 2004. Biosens. Bioelectron. 20 (6), 1111–1117. Rigby, G.P., Ahmed, S., Horseman, G., Vadgama, P., 1999. Anal. Chim. Acta 385 (1–3), 23–32. Shu, F.R., Wilson, G.S., 1976. Anal. Chem. 48 (12), 1679–1686. Sulak, M.T., Gokdogan, O., Gulce, A., Gulce, H., 2006. Biosens. Bioelectron. 21 (9), 1719–1726. Tao, W.Y., Liu, Y.J., Pan, D.W., Nie, L.H., Yao, S.Z., 2004. Bioelectrochemistry 65 (1), 51–58. Tian, Y., Wang, J.X., Wang, Z., Wang, S.C., 2005. Sens. Actuator B 104 (1), 23–28.

2905

Trojanowicz, M., Miernik, A., 2001. Electrochim. Acta 46 (7), 1053–1061. Vidal, J.C., Garcia-Ruiz, E., Espuelas, J., Aramendia, T., Castillo, J.R., 2003. Anal. Bioanal. Chem. 377 (2), 273–280. Wilson, G.S., Hu, Y.B., 2000. Chem. Rev. 100 (7), 2693–2704. Yang, H., Chung, T.D., Kim, Y.T., Choi, C.A., Jun, C.H., Kim, H.C., 2002. Biosens. Bioelectron. 17 (3), 251–259. Ye, J.S., Wen, Y., Zhang, W.D., Cui, H.F., Xu, G.Q., Sheu, F.S., 2005. Electroanalysis 17 (1), 89–96. Yoshimura, K., Hozumi, K., 1996. Microchem. J. 53 (4), 404–412. Zhang, Z.E., Liu, H.Y., Deng, J.Q., 1996. Anal. Chem. 68 (9), 1632–1638. Zhou, H.H., Chen, H., Luo, S.L., Chen, J.H., Wei, W.Z., Kuang, Y.F., 2005. Biosens. Bioelectron. 20 (7), 1305–1311.