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Sensors and Actuators B 243 (2017) 589–595

Contents lists available at ScienceDirect

Sensors and Actuators B: Chemical

jo ur nal home page: www.elsev ier .com/ locate /snb

abrication of a disposable non-enzymatic electrochemical creatinineensor

eethu Raveendrana,b, Resmi P.E. a, Ramachandran T.a, Bipin G. Nairc,.G. Satheesh Babua,∗

Department of Sciences, Amrita School of Engineering, Coimbatore, Amrita Vishwa Vidyapeetham, Amrita University, IndiaDepartment of ECE, Amrita School of Engineering, Coimbatore, Amrita Vishwa Vidyapeetham, Amrita University, IndiaAmrita School of Biotechnology, Amritapuri, Amrita Vishwa Vidyapeetham, Amrita University, India

r t i c l e i n f o

rticle history:eceived 6 August 2016eceived in revised form1 November 2016ccepted 29 November 2016

a b s t r a c t

A disposable non-enzymatic sensor for creatinine was developed by electrodepositing copper on screenprinted carbon electrodes. The sensor was characterized using electrochemical and microscopic tech-niques. Electrochemical detection of creatinine was carried out in phosphate buffer solution of pH 7.4. Theestimation was based on the formation of soluble copper-creatinine complex. The formation of copper-creatinine complex was established using the pseudoperoxidase activity of copper-creatinine complex.

vailable online 2 December 2016

eywords:on-enzymatic creatinine sensoropper electrodeopper-creatinine complex

The sensor showed a detection limit of 0.0746 �M with a linear range of 6–378 �M. The sensor exhibiteda stable response to creatinine and found to be free from interference from molecules like urea, glucose,ascorbic acid and dopamine. Real sample analysis was carried out with blood serum.

© 2016 Elsevier B.V. All rights reserved.

seudoperoxidase activity

. Introduction

Creatinine is a nitrogenous waste product formed from crea-ine and phosphocreatine in muscles and removed from the bodyy kidneys through glomerular filtration at a relatively constantate [1]. Normal range of creatinine is 40–150 �M in blood serumut can exceed 1000 �M in certain pathological conditions [2].lomerular filtration rate (GFR), is an important index of renal func-

ion and is inversely related to plasma creatinine [3]. Thereforeuantitative determination of creatinine in body fluids is used toecide adequacy of kidney function, the severity of kidney damagend the progression of kidney disease [4].

The conventional techniques for creatinine detection involveither electrochemical detection using three enzymes or opticaletection by Jaffe reaction [5]. The enzymatic detection comprisesf three enzymes, creatinine amidohydrolase, creatine amidohy-rolase and sarcosine oxidase [6]. This procedure involves the

mmobilization of enzymes on the electrode surface and theetection of H2O2 produced by the enzymatic reaction usingoltammetry or amperometry [7–12]. Different approaches for the

∗ Corresponding author.E-mail addresses: tgsatheesh@gmail.com, tg satheesh@cb.amrita.edu

T.G. Satheesh Babu).

ttp://dx.doi.org/10.1016/j.snb.2016.11.158925-4005/© 2016 Elsevier B.V. All rights reserved.

construction of enzymatic creatinine sensor have been reviewed[13–15]. Traditional enzymatic detection techniques are plaguedby poor storage stability, stringent operating conditions, interfer-ence from other species and high cost. Clinical analysis of creatinineis based on spectrophotometric determination by Jaffe reaction[16]. Chemicals in blood such as sugars, urea, uric acid, pyruvate,dopamine, acetone, acetoacetic acid, fructose, glucose, ascorbate,cefoxitin, cephalotin, cefatril and cefazolin interfere with this tech-nique [13]. Selectivity can be enhanced by the use of molecularlyimprinted polymers which can specifically accommodate targetanalyte through their cavities [17,18]. Khadro et al. developedmolecularly imprinted polymer based electrochemical sensor usingpolyethylene-co-vinyl alcohol [19]. But the dynamic range of sen-sor could not match with the normal creatinine range and elevatedranges during kidney failure.

Enzymeless impedemetric biosensor for creatinine was devel-oped using �-cyclodextrin and poly-3,4-ethylenedioxythiophenemodified glassy carbon electrode [20]. Though the sensor exhibitedfast response and good stability, it had poor selectivity and reduceddynamic range. Only a few reports are available on creatininesensors based on direct electrooxidation. Quantitative determina-

tion of creatinine was performed by preconcentration followed byoxidative stripping on screen printed carbon electrode [21].

Creatinine forms complexes with different transition metal ionssuch as Ag(I), Hg(II), Cd(II), Zn(II), Co(II), Ni(II), Cu(II), Pt(II) and

5 nd Actuators B 243 (2017) 589–595

P[toararffeot

eoihwtlo

2

2

di(cPw

2

etpwKrwtnes

2

pi(to

2

tt0c

Fig. 1. Electrochemical impedance spectra of SPCE (a) and Cu/SPCE (b) in

90 J. Raveendran et al. / Sensors a

d(II) since it has several donor groups in its main tautomeric form22]. This property has been utilized for the quantitative determina-ion of creatinine. Optical sensors utilizing the peroxidase activityf copper creatinine complex with tetramethylbenzidine (TMB)nd diisopropyl benzene dihydroperoxide (DBDH) [23] has beeneported. Copper-creatinine complex oxidizes paraphenylenedi-mine and the product couples with dimethylamino benzoicacidesulting in absorbance at 710 nm [24]. An amperometric methodor the determination of creatinine was developed based on theormation of a soluble copper–creatinine complex on the copperlectrode. Oxidative current during the regeneration of the surfacexide layer was monitored and it was correlated to the concentra-ion of creatinine [25].

In this paper we report the development and testing of a non-nzymatic creatinine sensor with high sensitivity and linear rangef detection. The sensor electrode was fabricated by electrodeposit-ng copper on screen printed carbon electrodes (SPCE). The sensoras been thoroughly characterized for its morphology and testedith physiological samples. Single step fabrication using cost effec-

ive and easily available materials, good sensitivity and excellentinear range of detection make the sensor a promising tool for pointf care testing (POCT).

. Experimental

.1. Chemicals

Creatinine, d-(+)-glucose, l-ascorbic acid, urea, uric acid andopamine were purchased from Sigma-Aldrich. All the other chem-

cals were of analytical grade and used as received. Phosphate bufferPB) was prepared by dissolving Na2HPO4 and KH2PO4 salts. Carbononducting paste (BQ242) was purchased from Du Pont Companyte. Ltd, Singapore. All solutions used in this study were preparedith millipore water (15.2 M� cm, Millipore, Germany).

.2. Instrumentation

Electrochemical experiments were carried out using CHI 660Clectrochemical workstation (CH Instruments, TX, USA). The elec-rochemical cell consisted of a three-electrode system with screenrinted carbon electrode (SPCE) as working electrode, a platinumire as counter electrode and saturated calomel (Hg/Hg2Cl2 (sat.Cl)) as reference electrode. All potentials used in this work are ineference to calomel electrode. Absorption spectra were recordedith UV–vis spectrophotometer (Pharmaspec 1700, Shimadzu) in

he range 200–800 nm. Surface morphology was studied with scan-ing electron microscope (SEM) (JEOL Model JSM-6390LV) andlemental analysis was carried out with Energy-dispersive X-raypectroscopy (EDS) spectrum (Hitachi SS6600).

.3. Fabrication of the sensor electrode

SPCEs were fabricated by screen printing carbon ink onolyethylene terephtalate (PET) substrate and cured for 2 h at 60 ◦C

n a hot air oven. The area of the working electrode is 3.14 mm2

r = 1 mm). Copper was electrodeposited on the SPCE from a solu-ion containing 0.1 M CuSO4 in 0.1 M H2SO4 at a constant potentialf −0.6 V versus calomel (Hg\Hg2Cl2) electrode.

.4. Characterization of bare and modified electrodes

Bare and modified electrodes were characterized by elec-

rochemical impedance spectroscopy (EIS) in a solution con-aining 5 mM K3[Fe(CN)6] in 1.5 M KCl at a frequency range of.01 Hz–100 kHz with an amplitude of 5 mV. EIS Experiments werearried out on the bare and modified electrodes at open circuit

5 mM K3[Fe(CN)6]/1.5 M KCl at open circuit potential under a frequency range of0.01 Hz to 100 kHz with an amplitude of 5 mV (Inset: Enlarged impedance spectraof Cu/SPCE).

potentials. Scanning electron microscopy was carried out at anenergy of 20 kV. Elemental analysis of bare and modified electrodeswas studied with EDS analysis.

2.5. Electrochemical detection of creatinine

Electrochemical detection of creatinine was carried out voltam-metrically. Cyclic voltammograms (CVs) were recorded at apotential window of −0.6 to 0.6 V at a scan rate of 100 mV/s.Creatinine solution was injected into the PB so the resultant con-centration varies in the physiological range. The interference ofglucose, ascorbic acid, urea, uric acid and dopamine was studiedby injecting the respective solutions into the PBS.

2.6. Preparation of samples for UV–vis analysis

A freshly polished metallic copper strip was selectively maskedusing teflon tape to expose an area of 1 cm × 2 cm and used as work-ing electrode. CV was carried out by cycling the potential between−0.6 to 0.6 V at a scan rate of 100 mV/s in PB of pH 7.4 contain-ing 5 mM creatinine with constant stirring. 50 �L of the solutionwas drawn at every 100 potential cycles and mixed with 50 �L ofTMB-H2O2 solution, incubated for 15 min at room temperature andsubjected to absorption spectroscopy.

2.7. Real sample analysis

Blood samples were collected from healthy volunteers and keptin refrigerator for eight hours to clot completely. Then the sam-ples were centrifuged at 12,000 rpm for 20 min. The sera carefullyseparated using pipette and tested. Prior to testing the serum wasincubated for 10 min with PbO2 powder (0.6 g for 4 mL) to removeinterference of uric acid.

3. Results and discussion

3.1. Electrochemical characterization of modified electrode

Electron transfer behavior of electrode-electrolyte interface wasstudied by electrochemical impedence sepectroscopy (EIS) and theresults obtained are presented as Nyquist plot. Semicircular region

at higher frequencies indicate kinetically controlled processes andthe semicircular diameter denotes charge transfer resistance, Rct.Kinetically sluggish systems were characterized with large Rct val-ues. The Nyquist plots (Fig. 1) obtained on bare (a) and copper

J. Raveendran et al. / Sensors and Actuators B 243 (2017) 589–595 591

Cu/SPCE (B) EDS spectrum of Cu/SPCE (C).

mmTlFsa

3

T(dt

sTm

3e

mv

Fig. 3. CVs recorded on bare and modified electrodes in 0.1 M PB with and without

Fig. 2. SEM images of SPCE (A) and

odified SPCE (b) in 5 mM K3[Fe(CN)6]/1.5 M KCl show that the for-er has very high charge transfer resistance about 75 k� (curve a).

he semicircular region is almost absent in curve b, indicating veryow charge transfer resistance of the copper modified electrode.rom this study it is obvious that the electrodeposition of copper oncreen printed electrode has increased the overall charge transfercross the electrode-electrolyte interface.

.2. Surface characterization

Fig. 2 shows the SEM images of bare and modified electrodes.he characteristic flake like structures were observed on SPCEFig. 2A) and electrodeposition of copper resulted in the uniformistribution of granular Cu mesoparticles of around 1 �m size onhe SPCE electrode (Fig. 2B).

From the EDS spectrum (Fig. 2C) it is evident that the electrodeurface contains large amount of Cu and very small amount of O.his supports the observation that the deposit exists mainly asetallic copper.

.3. Electrochemical behaviour of creatinine on modifiedlectrode

Fig. 3 shows the cyclic voltammograms obtained on bare andodified electrodes in PB with and without creatinine. The cyclic

oltammograms obtained on the modified electrode in PB solution

62.5 �M creatinine at a scan rate of 100 mV/s (inset: enlarged view of CVs obtainedon bare electrode).

show two anodic peaks and two cathodic peaks. The anodic peaksare attributed to the oxidation of copper and the cathodic peaks are

due to the reverse conversion of copper oxide to copper metal. Inthe presence of creatinine, the anodic peak observed at 0.2 V showsan increase in current whereas cathodic peak current decreases.CV obtained on bare SPCE does not show any faradic current in

592 J. Raveendran et al. / Sensors and Actuators B 243 (2017) 589–595

Fig. 4. CVs on Cu/SPCE in 0.1 M PB with a creatinine concentrations of 62.5 �M atd(

ttwcc

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i

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d

3c

ca

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been reported previously with other biomolecules also [29–31]. To

ifferent scan rates (a → k: 10, 20, 50, 70, 100, 150, 200, 300, 500, 700, 1000 mV/s)Inset: plot of peak current versus scan rate).

he presence of creatinine (Fig. 3 inset). It is proposed that duringhe forward scan, copper (II) ions on the electrode surface chelateith the creatinine and form soluble complex. The increase in peak

urrent can be attributed to the increased dissolution of copper asopper-creatinine complex.

The relationship between scan rate and peak current can be usedo explore valuable information regarding electrochemical behav-or of creatinine on modified electrode in PB. CVs were recorded on

odified electrode in PB pH 7.4 with 62.5 �M creatinine at differentcan rates from 10 to 1000 mV/s (Fig. 4). The increase in peak cur-ents was found to be directly proportional to increase in scan rateith a regression coefficient of R = 0.991. This indicates that thisernstian process is adsorption controlled and obeys the relation,

p = n2f 2

4RT�A�0

here ip, n, F, R,T, �, A and �o are the peak current, number oflectrons involved in the reaction, faraday’s constant, universal gasonstant, absolute temperature, scan rate, area of the electrode andmount of adsorbed species at time respectively. It was also seenhat the peak potential shifts with scan rate, which is normallybserved when the reactant is weakly adsorbed onto the electrodeurface [26].

This observation was further confirmed by well known pseu-operoxidase activity of the copper-creatinine complex [27].

.4. Spectroscopic evidence for the formation ofopper-creatinine complex

The UV–vis spectrum of the solution obtained after the potentialycling of Cu electrode in phosphate buffer containing creatinine,fter mixing with TMB-H2O2 is presented in Fig. 5.

The peaks observed at 375 nm and 652 nm are due to the onelectron oxidation of TMB and the resulted charge-transfer com-lex. The solution then turned to yellow (�max = 450 nm) is due tohe two electron oxidation product of TMB. Similar results wereeported by Josephy et al. for the oxidation of TMB using HRP [28].he control solutions did not show any characteristic peaks for TMBxidation. The increase in intensity of absorption peaks with num-er of potential cycles can be attributed to the formation of moremount of copper-creatinine complex in the solution and henceore TMB oxidation. No absorption peak was observed when the

est was performed after oxidizing the copper strip extensivelysing CV in PB solution in the absence of creatinine. From the aboveiscussion it is obvious that the absorption peak is due to the oxida-

Fig. 5. Absorption spectra obtained for TMB-H2O2 and 5 mM creatinine in PB after0, 100, 200, 300 and 400 potential cycles.

tion product and that is possible only due to the pseudoperoxidaseactivity of the copper-creatinine complex.

3.5. Effect of electrodeposition time and electrolyte concentrationon response current

In order to study the influence of the amount of Cu on theelectrode surface on response current, various electrodes werefabricated with different amounts of Cu by varying the deposi-tion parameters such as time of deposition and concentration ofelectrolyte solution. The electrodeposition time was varied from50 to 400 s and the sensitivity of the electrodes to creatinine wascompared. It is found that the sensor obtained after 100 s of elec-trodeposition showed maximum sensitivity and further increase indeposition time did not show any improvement in response, ratherit exhibited a shift in oxidation potentials to more positive values.Thus deposition for 100 s was considered as the optimum.

Similarly, electrodeposition was carried out at the optimizedpotential by varying the electrolyte concentration from 0.05 to 1 M.Maximum current response with creatinine was obtained for themodified electrode obtained in 0.1 M deposition bath. Hence all fur-ther electrodeposition was carried out in 0.1 M H2SO4 containing0.1 M CuSO4 at −0.6 V for 100 s.

3.6. Electrochemical detection of creatinine

The electrochemical detection of creatinine on the modifiedelectrode was carried out using CV in PB of pH 7.4. It was foundthat the peak current increases linearly with increase in the creati-nine concentration (Fig. 6). It is believed that on applying potentialthe copper undergoes electrochemical oxidation and the cupricions combine with the creatinine in the solution and form solu-ble copper-creatinine complex. Due to the dissolution of complexfresh copper surface will be exposed and on increasing the con-centration of creatinine more amount of Cu undergone oxidation.The sensor exhibited a linear response at lower concentrationsand found to deviate from linearity at higher concentrations (Fig.S1). Since the oxidation proceeds through adsorption, at lowerconcentrations the catalytic surface area of copper is sufficientfor complexation with creatinine and at higher concentrationsit is insufficient for complexation. This could be the reason forthe change in slope of the calibration curve. Similar results have

obtain a single linear response over the whole calibration rangelinearization of the x-axis was performed by taking the logarithmicscale of the concentration. The calibration equation obtained was IP

J. Raveendran et al. / Sensors and Actuators B 243 (2017) 589–595 593

eatinine at a scan rate of 100 mV/s (A), plot of peak current versus concentration (B).

(w

3

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3

twAwap(fsut(

Fig. 7. Linear sweep voltammograms at a scan rate of 100 mV/s on the modified

Fig. 6. CVs on modified electrode in 0.1 M PB with different concentrations of cr

�A) = 10.082log10C (�M) − 3.546. The limit of detection of sensoras statistically calculated (S/N = 3) as 0.0746 �M.

.7. Interference study

Interference of molecules such as glucose, ascorbic acid, urea,ric acid and dopamine was tested by injecting physiological con-entrations (glucose- 6 mM, urea- 6 mM, ascorbic acid- 125 �M,opamine-62.5 �M, uric acid- 125 �M) in the presence of 62.5 �Mreatinine. The modified electrode was found to be free from thenterference of these commonly interfering species except uriccid [Fig. S2]. Interference of uric acid was effectively removedy incubating the samples with PbO2 powder (0.6 g for 4 mL solu-ion) for 10 min [32,33] [Fig. S3]. Percentage response of interfering

olecules over the response of creatinine was provided in support-ng information (Table S1).

.8. Real sample analysis

Serum samples of known concentrations of creatinine wereested using the developed sensor and the results were comparedith that of the result obtained for creatinine solution in PB Fig. 7.s observed with creatinine samples (curve b), a well defined peakas observed with serum samples (curve c). It was found that on

dding serum a small shift from the baseline occurs. Inorder to com-ensate this shift, the calibration equation (IP (�A) = 10.082log10C�M) – 3.546) was slightly modified. The y intercept was increasedrom −3.546 to 0.09588 and the slope of the equation remained

ame. Using the new calibration curve, estimation of creatinine innknown blood samples was performed. The new calibration equa-ion for creatinine in the blood sample is IP (�A) = 10.082log10C�M) + 0.09588.

electrode in 0.1 M PB (a) with 6.25 �M creatinine (b), serum (c) and serum spikedwith 6.25 �M creatinine (d).

Using this calibration equation the estimation of creatinine inunknown blood samples was performed and compared with thatof the clinical method (Jaffe reaction). The percentage recoverywas calculated and found to be 98% and 95.6% for sample 1 and2 respectively Table 1. From the table it is clear that the resultsobtained from fabricated sensor is in line with those obtained fromconventional detection method.

Test was conducted by spiking known creatinine concentrationinto serum samples. It was found that the calculated concentra-

tion is almost equal to the estimated creatinine concentration afterspiking with 6.25 �M creatinine, with a recovery of 100.725% (TableS2).

594 J. Raveendran et al. / Sensors and Actuators B 243 (2017) 589–595

Fig. 8. Peak current obtained on an electrode for 10 potential cycles (A) an

Table 1Creatinine concentration present in the serum samples determined by the clinicallaboratory method and the sensor developed in this study.

Clinical laboratoryanalysis

Cu/SPCE sensor Recovery (%)

Sample 1 59.22 �M 58.06 �M 98.04

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[22] M. Mitewa, Coordination properties of the bioligands creatinine and creatinein various reaction media, Coord. Chem. Rev. 140 (1995) 1–25.

Sample 2 54.808 �M 52.40 �M 95.60

.9. Repeatability and reproducibility

The operational stability of the sensor was tested by exam-ning the performance of a single electrode repeatedly in the PBolution containing 62.5 �M creatinine. It was found that for morehan 10 trials the electrode exhibited similar response (RSD = 4.1%,D = 0.517, n = 10) [Fig. 8A].

The reproducibility of the modified electrode was analyzed byomparing the response current of 10 different sensor electrodeso 62.5 �M of creatinine. All the sensors exhibited similar responseurrent and a variation of only 4.8% was observed [Fig. 8B]. Thisndicates that the fabrication process optimized is highly reliablend uniform.

. Conclusion

A disposable non-enzymatic, electrochemical creatinine sensoras developed by the electrodeposition of copper on screen printed

arbon electrodes. Detection was based on the formation of solu-le copper–creatinine complex formation, which was establishedy the pseudoperoxidase activity of copper–creatinine complex.he sensor exhibited excellent catalytic response towards creati-ine in the range of 6.25–378.5 �M at the physiological pH making

t highly suitable for real sample analysis. Interference from otherpecies was found to be negligible or minimal. Comparison of blooderum analysis results with those obtained from conventional clin-cal methods was highly correlative which further reiterates theotential of the developed sensor for POCT application.

cknowledgements

Authors greatly acknowledge the Department of Biotechnol-gy (DBT), Government of India for the financial support underhe projects (Project No BT/PR4076/MED/32/221/2011). Jeethu

aveendran express her sincere thanks to Council of Scientific and

ndustrial Research (CSIR) for the financial support under CSIR-SRFirect scheme (Sanction No. 09/1034(0003)/2K15-EMR-I).

[

d on 10 different electrodes (B) in PB containing 62.5 �M creatinine.

Appendix A. Supplementary data

Supplementary data associated with this article can be found, inthe online version, at http://dx.doi.org/10.1016/j.snb.2016.11.158.

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uators B 243 (2017) 589–595 595

Biographies

Ms. Jeethu Raveendran received her B.E in Biomedical Engineering from AnnaUniversity, Chennai and M.Tech in Biomedical Engineering from Amrita VishwaVidyapeetham University, Coimbatore. Currently she is a PhD student in AmritaUniversity. Her research interests focuses on the development of electrochemicalbiosensors, microfluidics and biosignal processing.

Ms. Resmi P.E. received B. Sc degree in chemistry from Amrita School of Art andScience, Mysore and M.Sc degree from Amrita School of Art and Science, Kollam.Now she is doing her PhD in Amrita University. Her research focuses on developmentof non-enzymatic biosensors.

Dr. T. Ramachandran received B.Sc, M.Sc and PhD degree from Madras Univer-sity. He had post doctoral research at Georgetown University (Washington, USA).Presently he is a Professor in Amrita University. His research interests are industrialelectrochemical processes, fuel cells and biosensors.

Dr. Bipin G. Nair received B.Sc degree from Gujarat University, M.Sc and PhD degreefrom Maharaja Sayajirao University of Baroda. He had four years of Post doctoralresearch in University of Tennessee, Memphis. Now he is Professor and Dean ofSchool of Biotechnology, Amrita University, Kollam.

synthesis of nanomaterials, biosensors and lab on a chip.