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Current Applied Physics 12 (2012) 1118e1124

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Current Applied Physics

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Novel amperometric glucose biosensor based on covalent immobilizationof glucose oxidase on poly(pyrrole propylic acid)/Au nanocomposite

Mehmet Senel*, Cevdet Nergiz*

Department of Chemistry, Faculty of Arts and Sciences, Fatih University, Buyukcekmece Kampusu, B.Cekmece, Istanbul 34500, Turkey

a r t i c l e i n f o

Article history:Received 4 January 2012Received in revised form1 February 2012Accepted 2 February 2012Available online 20 February 2012

Keywords:BiosensorGlucose oxidasePyrroleImmobilizationNanocomposite

* Corresponding authors. Tel.: þ902128663300; faxE-mail addresses:msenel@fatih.edu.tr (M. Senel), c

1567-1739/$ e see front matter � 2012 Elsevier B.V.doi:10.1016/j.cap.2012.02.004

a b s t r a c t

A glucose biosensor is fabricated with immobilization of glucose oxidase (GOx) on poly(pyrrole propylicacid)/Au nanocomposite by covalent attachment. Poly(pyrrole propylic acid) (PPyAA)/Au nanocompositewas prepared by chemical oxidation of pyrrole propylic acid monomer by using chloroauric acid (HAuCl4)as an oxidizing agent. The obtained nanocomposites were used to fabricate highly sensitive ampero-metric glucose biosensor which exhibited a high and reproducible sensitivity of 0.42 mA/mM, responsetime w2 s, linear dynamic range from 1 to 18 mM, correlation coefficient of R2¼ 0.9981, and limit ofdetection (LOD), based on S/N ratio (S/N¼ 3) of 0.05 mM. A value of 1.83 mM for the apparent MichaeliseMenten constant Kapp

m was obtained. The high sensitivity, wider linear range, good reproducibility andstability make this biosensor a promising candidate for portable amperometric glucose biosensor.

� 2012 Elsevier B.V. All rights reserved.

1. Introduction

To obtain efficient enzyme electrodes, the researchers haveattempted to immobilize enzymes on the surfaces of the conduct-ing polymer films. Polypyrrole (PPy), as one of the most importantconducting polymer, has been widely studied in electrochemicalbiosensor applications in recent years due to its stability, conduc-tivity and biocompatibility [1]. PPy can be synthesized by eitherchemical or electrochemical methods. The effects of substratematerials, synthesis methods, and the dopands on the properties ofthe PPy have been studied [2e8]. Some techniques, includingentrapment [9e11], adsorption [12,13], affinity interaction [14,15]and covalent immobilization [16e18] have been successfully con-ducted for the effective enzyme immobilization. The entrapment ofbiomolecules a simple one-stepmethod during the electrochemicalpolymerization of pyrrole, but it suffers greatly from the pooraccessibility of the target molecules due to its hydrophobicity. Thecovalent immobilization of the enzyme on the surface of the filminvolves the attachment of enzyme molecules through chemicalbinding between enzyme and the carboxyl or amine groups on thesurface. This functional group of the polymer surface introducedeither by the post-functionalization of the PPy film or the initial

: þ902128663402.nergiz@fatih.edu.tr (C.Nergiz).

All rights reserved.

polymerization of the functionalized pyrrole. This approach allowsbetter orientation for higher activity and long-time stability ofenzymes, making it a preferable method for construction of theenzyme electrode. In a previous study, electrochemically synthe-sized conducting co-polymer of the pyrrole and pyrrole propylicacid films were used to covalently immobilize glucose oxidase bycondensation reaction with COOH groups on the films [19].According to this study, the amount of pyrrole propylic acid (PyAA)monomer in the co-polymer composition affects the performanceof the films in amperometric biosensing. Higher amount of thePyAAmonomer in the co-polymermay decrease the conductivity ofthe film compared with the homopolymer of pyrrole.

Due to their unique properties such as catalytic activities, opticalproperties, and biocompatibility gold nanoparticles have beenextensively utilized in recent years. Several researches have beendevoted to fabricate gold nanoparticles dispersed onto varioussubstrates such as carbon paste electrode, self-assembled mono-layer, conducting and non-conducting polymers [20e23]. It is wellknown that the PPy and AuNPs bear good electrochemical behaviorandhavebeenwidely used in thefield of electro catalysis and electroanalysis [24]. Considering the individual advantages of the PPyAAand the Au nanoparticles, the PPyAAeAu nanocomposite designedin this work for highly sensitive and stable amperometric biosensor,in which the sensing film could benefit from the good conductivity,the enhanced hydrophilicity and covalent binding ability of PPyAA,and the superior stability and conductivity of Au nanoparticles.

M. Senel, C. Nergiz / Current Applied Physics 12 (2012) 1118e1124 1119

Determination of glucose has an importance in biological fluidssuch as blood and urine to make a diagnosis and treatment ofdiabetes mellitus. Most of the amperometric glucose biosensors arebased on the specific recognition of glucose by the enzyme glucoseoxidase. Glucose oxidase contains two tightly bound flavin adeninedinucleotide (FAD) cofactors and catalyzes the oxidation of b-D-glucose to gluconic acid, by utilizing molecular oxygen as an elec-tron acceptor with simultaneous production hydrogen peroxide.The following reactions display the transducing principle in Fig. 1[25].

The quantification of glucose can be achieved via electro-chemical detection of the enzymatically liberated H2O2.

The aim of this study is to construct a novel efficient glucosebiosensor based on glucose oxidase enzyme. Herein, we usedchemically synthesized poly(pyrrole propylic acid)/Au nano-composite to fabricate a novel glucose biosensor. GOx was cova-lently immobilized on the surface of the nanocomposite filmthrough amide linkages by condensation reaction between NH2groups of the GOx and COOH groups of the nanocomposite film.The performance of the biosensor was studied in detail.

2. Experimental

2.1. Materials and apparatus

Glucose Oxidase (GOx) (EC 1.1.3.4), 1-Cyclohexyl-3-(2-morpholinoethyl)-carbodiimide metho-p-toluenesulfonate (CMC)and 1-(2-cyanoethyl)pyrrole (Py-CN) were obtained from AldrichChemical Co. All other chemicals were of analytical grade and usedwithout further purification.

The FT-IR-ATR spectra (4000e400 cm�1) were recorded witha Bruker spectrometer. NMR spectra were recorded in CDCl3 usinga Bruker 400 MHz spectrometer.

X-ray powder diffraction (XRD) analysis was conducted ona Rigaku SmartLab Diffractometer operated at 40 kV and 35 mAusing Cu Ka radiation.

Electrochemical impedance spectroscopy (EIS) measurementswere carried out using a CHI Model 6005 electrochemical analyzerin a background solution of 5 mM Fe3þ/Fe2þ phosphate buffer (pH7.0) at a normal potential. The alternating voltage was 5 mV and thefrequency range was 5.0�10�2e1.0�106 Hz.

Electrochemical polymerizations and measurements were per-formed using a CHI Model 842B electrochemical analyzer. A smallglassy carbonworking electrode (GCE) (2 mmdiameter), a platinumwire counter electrode (0.2 mm diameter), an Ag/AgCl-saturated

Fig. 1. Reaction of glucose

KCl reference electrode, and a conventional three-electrode elec-trochemical cell were purchased from CH Instruments.

2.2. Synthesis of 1-(2-carboxyethyl)pyrrole (PyAA)

1-(2-carboxyethyl)pyrrole was obtained by hydrolysis of 1-(2-cyanoethyl)pyrrole (Py-CN) according to literature [26]: a mixtureof 25 g of Py-CN and 100 ml of 15% potassium hydroxide solutionwas stirred at 50 �C for 40 h. Then the mixture was cooled to roomtemperature and acidified by hydrochloric acid. After extractionwith ether, the crude product (colorless crystals) was collected onevaporation of ether. The crude product was dissolved in ether andpurified by recrystallization from the ether solution. The productwas identified as 1-(2-carboxyethyl)pyrrole by means of FT-IR-ATR,1H, and 13C NMR spectroscopy (not shown).

2.3. Electrode preparation and immobilization of GOx onPPyAAeAu nanocomposite film

Before modification, the GCE working electrode was polishedwith 0.3 mm alumina powder and sonicated for at least 10 min indistilled water. Then, the electrode was rinsed alternatively withmethanol and water and dried under a nitrogen stream. Synthesisof PPyAAeAu nanocomposite and subsequent modification of theGCE electrode was performed by depositing 2.5 ml of 0.0125 MHAuCl4 and 2.5 ml of 0.0925 M PyAA monomer solutions ontothe electrode surface and allowing to react and dry for 30 min.The electrode was then thoroughly rinsed with distilled waterbefore use.

The nanocomposite film-coated electrode was immersed in3.0 ml distilled water containing 25 mg GOx and 120 mg CMC, andallowed to remain for 24 h at 4 �C. The co-polymer films treatedthus, i.e. GOx electrodes, were then rinsed with distilled water andstored in 0.1 M phosphate buffer (pH 7.4) at 4 �C.

2.4. Amperometric measurements

All amperometric measurements were carried out at roomtemperature in stirred solutions by applying the desired potentialand allowing the steady state current to be reached. Once prepared,the GOx electrodes were immersed in 10 ml of 10 mM PBS (phos-phate buffer saline) solution at pH 7.5 and the amperometricresponse to the addition of a known amount of glucose solutionwas recorded. Data collected with freshly prepared enzyme elec-trodes refer to the average of three experiments.

oxidase with glucose.

M. Senel, C. Nergiz / Current Applied Physics 12 (2012) 1118e11241120

3. Results and discussion

3.1. Fabrication of the enzyme electrode

The GOx immobilized electrode was constructed within two-step as illustrated in Fig. 2. In first step, the PPyAAeAu polymericnanostructure was prepared directly onto the GCE surface ina single phase reaction between PyAA and HAuCl4 according toliterature [27]. In this process, chemical polymerization was doneby using HAuCl4 as an oxidizing agent and form uniformlydistributed gold nanoparticles within the polymeric structure. Innext step, PPyAAeAu nanocomposite coated electrode wasimmersed in the enzyme solutionwith containing CMC as couplingagent for one day to covalently immobilize enzyme.

The electrochemical behavior of PPyAAeAu and its comparisonwith PPyAA and bare electrode were investigated by cyclic vol-tammetry (Fig. 3A). A potential range of �0.2 to 0.7 V versus Ag/AgCl at a scan rate 100 mV/s was applied in the solution of 10 mMPBS solution, pH 7.5 containing [Fe(CN)6]3þ/4þ (5 mM). A charac-teristic reversible redox peaks of the Fe(II)/Fe(III) with an increasedoxidation/reduction current for the PPyAAeAu was observed ascompared for the PPyAA coated electrode. This enhanced electro-chemical behavior is attributed to the presence of gold nano-particles in the polymeric film [27,28]. For further characterization,cyclic voltammograms of (a) bare GCE, (b) GCEePPyAA and (c)GCEePPyAAeAu electrodes in 100 mM PB solution (pH 6.5) wereperformed (Fig. 3B). An oxidation peak at 1.23 V was shown for theGCEePPyAAeAu electrode and attributed to the oxidation of Au0 toAu3þ in the form of Au2O3 [29]. This result confirmed that thepresence of gold nanoparticles in the PPyAA film deposited on thesurface of the electrode.

Fig. 4 shows a typical Nyquist plot of electrochemical impedancespectroscopy (EIS) measurements of bare GCE, PPyAA andPPyAAeAu. The spectrum is a plot of imaginary part (Z00) versus realpart (Z0) of the impedance measured as a function of frequency u

which increases from right to left. The electron transfer resistance(Ret) was extracted parameters by fitting the EIS data into a suitableequivalent circuit on each independently fabricated electrode. TheRet can be used to describe the interface properties of the electrode.At a bare GCE, the redox process of the probe showed a low electrontransfer resistance 245 U (Fig. 4a). When PPyAA was casted on thebare electrode, the semicircle increased dramatically (Fig. 4b). The

Fig. 2. Schematic representation of the immobilizat

Ret value was about 2550 U, suggesting that the film of the polymerlayer on the electrode surface inhibited the interfacial electrontransfer. When Au nanoparticles containing polymer film wasprepared on the bare electrode surface, a barrier of the polymerfilm would increase the electron transfer efficiency and decreasethe Ret value to 1770 U (Fig. 4c).

Fig. 5 shows the FT-IR spectra of the pure PPyAA and PPyAAeAunanocomposite. As shown in Fig. 5A strong peak at 1710 cm�1

belongs to the C^O vibration of carboxylic acid substitution ofpyrrole ring. The peaks at 598 and 1396 cm�1 correspond to thepyrrole rings vibration, the bands at 1290, 1178 and 1034 cm�1 and918 and 780 cm�1 were assigned to ^CeH in-plane vibration andout-of-plane vibration, respectively [30]. As shown in Fig. 4b,formation of the gold nanoparticles led to some peaks of purePPyAA shifted to higher wave number. Overall, these two FT-IRspectra provided supportive evidence that PPyAAeAu nano-composite have been successfully prepared.

Fig. 6 shows the typical X-ray diffraction patterns of PPyAA andPPyAAeAu. In the XRD curve of the PPyAA and PPyAAeAu, a broadpeak centered at 2q¼ 22.5� related to the amorphous PPyAA wasobserved, which is similar to the conventional PPy [31]. The XRDcurve of PPyAAeAu shows, beside the PPyAA peak, four strongbands appeared with maximum intensity at 38.2�, 44.3�, 64.7� and77.8� due to the presence of Au nanoparticles. Those peaks can beassigned to the (111), (200), (220) and (311) lattice planes of fcc-Au.As expected due to the higher scattering intensity of the goldnanoparticles compared to the amorphous polymer structure, thepeaks assigned to the Au crystal plains appear with a higherintensity than the bands of polymer.

3.2. Optimization of experimental variables of biosensor

The optimization of the applied potential at the working elec-trode is fundamental to achieve the lowest detection limit and tominimize the electrochemical interfering species. The effect of theapplied potential on the current response to glucose has beeninvestigated by the addition of 10 mM glucose at various levels ofthe applied potential. Fig. 7A shows the relationship between theresponse current of the fabricated biosensor and the appliedvoltage. It can be seen that the current increases very slowlywhen the voltage is below þ0.4 V. However, if the voltageexcesses þ0.4 V, the current first rises drastically and then shows

ion of GOx on PPyAAeAu nanocomposite film.

Fig. 3. (A) Cyclic voltammograms of (a) bare GCE, (b) GCEePPyAA and (c) GCEeP-PyAAeAu electrodes in 10 mM PBS solution, pH 7.5 containing [Fe(CN)6]3�/4� (5 mM).(B) Cyclic voltammograms of (a) bare GCE, (b) GCEePPyAA and (c) GCEePPyAAeAuelectrodes in 100 mM PB solution, pH 7.5.

Fig. 4. Electrochemical impedance spectra of (a) bare GCE, (b) GCEePPyAA and (c) GCEePPyAAeAu electrodes in 10 mM PBS solution, pH 7.5 containing [Fe(CN)6]3�/4� (5 mM).

Fig. 5. FT-IR-ATR spectra of PPyAA and PPyAAeAu.

M. Senel, C. Nergiz / Current Applied Physics 12 (2012) 1118e1124 1121

the saturation behavior when voltage reaches þ0.6 V. So, þ0.6 Vwas chosen as an applied potential value for subsequentexperiments.

The pH effect on the amperometric biosensor performance wasinvestigated by measuring the current response to 10 mM glucoseat þ0.6 V. The effect of the pH on the enzyme electrode was pre-sented in Fig. 7B between pH 5.0 and 9.0. As seen in Fig. 7B, thebiosensor shows an optimal response at pH 7.5. So pH 7.5 of PBS isused as support electrolyte for glucose detection in further exper-iments, which is similar to that reported previously [32].

The effect of various temperatures on the biosensor perfor-mance was also studied between 25 and 60 �C. The activity of theenzyme electrode has been investigated by adding 10 mM glucosesolution in PBS, pH 7.5 solution by amperometric measurements,shown in Fig. 7C. It was observed that the response increased withtemperature increase, reaching a maximum at 45 �C, and thendecreased. Although the biosensor shows a maximum response at

Fig. 6. X-ray diffraction patterns of PPyAA and PPyAAeAu.

Fig. 7. (A) Effect of potential on the current response of the GOx immobilized elec-trodes by addition of 10 mM glucose at various levels of applied potential in 10 M PBS(pH 7.5 and w25 �C), (B) Effect of pH on GOx immobilized electrodes in 10 M PBS(w25 �C) and (C) Effect of temperature on GOx immobilized electrodes in 10 M PBS(pH 7.5).

Fig. 8. Amperometric response of enzyme electrodes to successive glucose injectionsat an applied potential þ0.6 V in stirred 10 mM PBS. The inset shows the calibrationcurve.

M. Senel, C. Nergiz / Current Applied Physics 12 (2012) 1118e11241122

45 �C, room temperature (25 �C) was chosen as working tempera-ture due to avoid possible solvent evaporation and ease of opera-tions. The dependence of amperometric current on temperature inan initial region can be expressed as an Arrhenius relationship

í(T0)¼ í0exp{�Ea/RT}

where í0 represents a collection of currents, R is the gas constant,T is the temperature in K degrees, and Ea is the activation energy.The activation energy for enzymatic reaction is calculated to be1.71 kJmol�1 from the slope of I� 1/T in the adoptive region oftemperature (Inset of Fig. 7C).

3.3. Amperometric response of the GOx electrode

Glucose oxidase converts glucose to D-glucono-1,5-lactoneenzymatically. The enzymatic reaction of glucose may be written asfollows:

glucoseþ O2/GOx

gluconic acidþH2O2;

H2O2/0:6V

O2 þ 2Hþ þ 2e�:

The hydrogen peroxide generated is detected at the workingelectrode at þ0.6 V versus Ag/AgCl electrode. Fig. 8 shows theconstant potential currentetime response and calibration curvegraphs of the glucose biosensor utilizing both without Au

Table 1Comparison of the analytical performances of the glucose biosensors.

Electrode RT(s)

Linear range(mM)

Detectionlimit (mM)

Sensitivity(mAmM�1 cm�2)

Ref.

PyCO2H/Au 2 1.0e18.0 0.05 0.42 This workPyCO2HcoPyFc e 1.0e5.0 6.9� 10�3 1.796 [37]PyCO2H e 1.0e80 e 0.85 [38]Py-Fc 4 1.0e18 e 0.05 [39]Py 15 0e5.0 e 8� 10�2 [40]Py e 1.0e22 e 0.27 [41]AuNP-Eggshell 30 8.33� 10�3e0.966 3.5� 10�3 e [42]

RT, response time; Py, poly(pyrrole); PyCO2H, poly (3-(1H-pyrrol-1-yl)propanoicacid); Au, Gold nanoparticle.

Table 2Glucose content determination in serum samples.

Samples Determinedby hospital (mM)

Measured bybiosensor (mM)a

RSD (%) Relativeerror (%)

1 5.51 5.35� 0.06 2.3 �2.902 4.85 5.05� 0.1 3.5 þ4.123 5.24 5.38� 0.08 3.2 þ2.67

a Average of the three measurements.

M. Senel, C. Nergiz / Current Applied Physics 12 (2012) 1118e1124 1123

nanoparticles and with Au nanoparticles for successive additions ofthe same amount of glucose. As is apparent from Fig. 8, theresponse of the electrode (PPyAAeAu-GOx) was w4 times as largeas that of the electrode (PPyAAeGOx). The result clearly demon-strates the activity of enzyme immobilized by the Au nanoparticleswas superior to the enzyme immobilized through PPyAA covalentbonding. The catalytic current of the enzyme electrodes rises withincreasing glucose concentration. It is likely that this behaviorfollows MichaeliseMenten kinetics, as observed for many amper-ometric enzyme electrodes [33e35]. The linear range can beobserved to be up to 16 mM with correlation coefficient (R) of0.9981, and then a plateau is reached gradually at the higherglucose concentration. The biosensor has a good detection limit of0.05 mM (signal-to-noise¼ 3), a high sensitivity of 0.42 mA/mMand a short response time (within w2 s).

The MichaeliseMenten analyses (Kappm and imax) represent the

linear response range and the dynamic range of the sensors,respectively. The MichaeliseMenten constant ðKapp

m Þ, determined

Fig. 9. (A) Long-time stability of GOx immobilized electrode. The amperometricresponses of these enzyme electrodes are regularly checked during 50 days by adding4 mM glucose solution (pH 7.5; w25 �C), (B) Currentetime response curve for theenzyme electrode obtained on subsequent additions of 0.1 mM ascorbic acid (AsA),0.3 mM acetaminophen (AA), 0.4 mM uric acid (UA) and 1 mM glucose (Gl).

from Fig. 8 using an electrochemical LineweavereBurk plot, was1.83 mM. This value is much smaller than in earlier studies [17,36],indicating that the present electrode exhibits a higher affinity forglucose. In Table 1 the analytical performance of this biosensor iscompared with that of the prior studies. The analytical character-istics of the enzyme electrode indicate that the enzyme waseffectively attached into the conducting polymer film and that thebioactivity of the immobilized enzyme was well preserved in theenzyme electrode.

3.4. Stability, reusability, interference and sample analysis studies

The long-term stability of enzyme electrode was evaluated bymeasuring its performance every few days. It can be seen fromFig. 9A the biosensor retains its activity after 10 days. After 50 days,the biosensor still active and retains around 50% of initial response,indicating that the immobilized GOx can efficiently retain itsbioactivity for quite long time.

In addition, the influences of common interfering electroactivesubstances in the analysis of plasma samples were studied andpresented in Fig. 9B. The interference tests were carried with threecommon interfering compounds: ascorbic acid, acetaminophenand uric acid, respectively. The interfering effects of adding 0.4 mMUA, 0.3 mM AA and 0.1 mM AsA are nearly eliminated at theenzyme electrode, compared to the high sensitivity of the resultglucose biosensor, the interference current can be negligible similaras the previous studies [42].

The practical use of the prepared biosensor was demonstratedby using human plasma samples. First, fresh plasma samples wereanalyzed in the local hospital, then; the samples were re-assayedwith the enzyme electrode. A plasma sample (1.0 ml) was addedinto 5 ml PBS (pH 7.5), and the response was obtained at constantpotential (þ0.6 V). The glucose content of the blood plasma canthen be calculated from the calibration plot of the enzyme elec-trode. The results in Table 2 are satisfactory and agree closely withthose by the biochemical analyzer in hospital.

4. Conclusions

The present study demonstrated the feasibility of developmentof a conducting poly(pyrrole propylic acid)/Au nanocompositebased amperometric biosensor for glucose determination. Theconducting polymer having a carboxylic acid functional group wasprepared via chemical oxidation by using HAuCl4 as an oxidizingagent. This functional polymer/nanocomposite is an efficient plat-form to immobilize enzyme and produce the enzyme electrode toexhibit a good performance in terms of dynamic range of detection,short response time and long lifetime and stability. Most healthyadults maintain fasting glucose levels above 4.0 mM (72 mg/dl),and develop symptoms of hypoglycemia when the glucose fallsbelow 4 mmol/L. A subject with a consistent range above 7 mM isgenerally held to have hyperglycemia, whereas a consistent rangebelow 4 mM is considered hypoglycemic. The results showed thatthe prepared new biosensor is suitable for measurement of theamount of glucose in a range 1.0e18.0 mM in biological fluids to

M. Senel, C. Nergiz / Current Applied Physics 12 (2012) 1118e11241124

make a diagnosis of diabetes. The simple method of fabrication ofpoly(pyrrole propylic acid)/Au nanocomposite/GOx electrode is anadditional advantages as compared with conventional methods.

Acknowledgments

This research was supported by grants from T.R. Prime MinistryState Planning Organization.

References

[1] W. Chen, C.M. Li, X. Yang, C.Q. Sun, C. Gao, Z.X. Zeng, J. Sawyer, Front. Biosci. 10(2005) 2518e2526.

[2] W. Hu, C.M. Li, X. Cui, H. Dong, Q. Zhou, Langmuir 23 (2007) 2761e2767.[3] D. Zane, G.B. Appetecchi, C. Bianchini, S. Passerini, A. Curulli, Electroanalysis 23

(2011) 1134e1141.[4] S. Sönmezo�glu, C.B. Durmus, R. Tas, G. Çankaya, M. Can, Semicond. Sci.

Technol. 26 (2011) 055011.[5] L. Tian, J. Qiu, Y.-C. Zhou, S.-G. Sun, Microchim. Acta 169 (2010) 269e275.[6] Viorel Brânzoi, Luisa Pilan, Florina Brânzoi, Reveu Roumaine De Chimie 54

(2009) 783e789.[7] Fatma Arslan, Sensors 8 (2008) 5492e5500.[8] Kh. Ghanbari, S.Z. Bathaie, M.F. Mousavi, Biosens. Bioelectron. 23 (2008)

1825e1831.[9] S. Akgöl, A. Denizli, J. Mol. Catal. B 28 (2004) 7e14.

[10] M.R. Guascito, D. Chirizzi, C.Malitesta, E.Mazzotta, Analyst 136 (2011) 164e173.[11] C. Malitesta, M.R. Guascito, Biosens. Bioelectron. 20 (2005) 1643e1647.[12] M. Senel, E. Çevik, M.F. Abasıyanık, A. Bozkurt, J. Appl. Polym. Sci. 119 (2011)

1931e1939.[13] M.C. Tsai, Y.C. Tsai, Sensor. Actuator. B Chem. 141 (2009) 592e598.[14] R.E. Ionescu, C. Gondran, I.A. Gheber, S. Cosnier, R.S. Marks, Anal. Chem. 76

(2004) 6808e6813.[15] X.Ren,D.Chen,X.Meng,F. Tang,A.Du, L.Zhang,Colloid.Surf.B72 (2009)188e192.

[16] M. Senel, M.F. Abasıyanık, Electroanalysis 22 (2010) 1765e1771.[17] M.F. Abasıyanık, M. Senel, J. Electroanal. Chem. 639 (2010) 21e26.[18] C. Ozdemir, F. Yeni, D. Odaci, S. Timur, Food Chem. 119 (2010) 380e385.[19] M. Shimomura, R. Miyata, T. Kuwahara, K. Oshima, S. Miyauchi, Eur. Polym. J.

43 (2007) 388e394.[20] A. Tkachenko, H. Xie, D. Coleman, D. Glom, J. Ryan, M. Anderson, S. Franzen,

D. Feldheim, J. Am. Chem. Soc. 125 (2003) 4700e4701.[21] N. Rosi, C. Mirkin, Chem. Rev. 105 (2005) 1547e1562.[22] M. Daniel, D. Astruc, Chem. Rev. 104 (2004) 293e346.[23] M.M. Ayad, J. Mater. Sci. 44 (2009) 6392e6397.[24] H. Zhang, X. Zhong, J.J. Xu, H.Y. Chen, Langmuir 24 (2008) 13748e13752.[25] S.B. Bankar, M.V. Bule, R.S. Singhal, L. Ananthanaryan, Biotechnol. Adv. 27

(2009) 489e501.[26] S.E. Wolowacz, B.F.Y. Yon-Hin, C.R. Lowe, Anal. Chem. 64 (1992) 1541e1545.[27] J. Njagi, S. Andreescu, Biosens. Bioelectron. 23 (2007) 168e175.[28] J. Kim, J.W. Grate, P. Wang, Chem. Eng. Sci. 61 (2006) 1017e1026.[29] L.D. Burke, P.F. Nugent, Gold Bull. 30 (1997) 43e53.[30] W. Chen, X.Y. Li, G. Xue, Z.Q. Wang, W.Q. Zou, Appl. Surf. Sci. 218 (2003)

215e218.[31] L.F. Warran, D.P. Anderson, J. Electrochem. Soc. 134 (1987) 101.[32] V.K. Gade, D.J. Shirale, P.D. Gaikwad, P.A. Savale, K.P. Kakde, H.J. Kharat,

M.D. Shirsat, React. Funct. Polym. 66 (2006) 1420e1426.[33] Y. Oztekin, A. Ramanaviciene, Z. Yazicigil, A.O. Solak, A. Ramanavicius, Biosens.

Bioelectron. 26 (2011) 2541e2546.[34] S. Tuncagil, S. Varis, L. Toppare, J. Mol. Catal. B 64 (2010) 195e199.[35] M.T. Sulak, Ö. Gökdo�gan, A. Gülce, H. Gülce, Biosens. Bioelectron. 21 (2006)

1719e1726.[36] A. Gülce, H. Gülce, J. Biochem. Biophys. Methods 62 (2005) 81e92.[37] M. Senel, Synthetic Met. 161 (2011) 1861e1868.[38] K. Kojima, T. Yamauchi, M. Shimomure, S. Miyauchi, Polymer 39 (1998)

2079e2082.[39] N.C. Foulds, C.R. Lowe, Anal. Chem. 60 (1988) 2473e2478.[40] L. Liu, N. Jia, Q. Zhou, M. Yan, Z. Jiang, Mater. Sci. Eng. C 27 (2007) 57e60.[41] W.J. Sung, Y.H. Bae, Biosens. Bioelectron. 18 (2003) 1231e1239.[42] B. Zheng, S. Xie, L. Qian, H. Yuan, D. Xiao, M.M.F. Choi, Sensor. Actuator. B

Chem. 152 (2011) 49e55.