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A highly sensitive electrochemical glucose sensor structuring with nickel hydroxide and enzyme glucose oxidase
Electrochimica Acta 108 (2013) 274–280
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A highly sensitive electrochemical glucose sensor structuring with nickel hydroxide and enzyme glucose oxidase Manjusha Mathew, N. Sandhyarani ∗ Nanoscience Research Laboratory, School of Nano Science and Technology, National Institute of Technology Calicut, Calicut, Kerala, India
a r t i c l e
i n f o
Article history: Received 11 May 2013 Received in revised form 27 June 2013 Accepted 4 July 2013 Available online xxx Keywords: Electrochemical sensor Nickel hydroxide Chitosan gold nanocomposite Cyclic voltammetry Glucose oxidase
a b s t r a c t A multilayered glucose biosensor with enhanced electron transport was fabricated via the sequential electrodeposition of chitosan gold nanocomposite (CGNC) and nickel hydroxide (Ni(OH)2 ) on a bare gold electrode and subsequent immobilization of glucose oxidase. A thin film of Ni(OH)2 deposited on CGNC modified gold electrode serves as an electrochemical redox probe as well as a matrix for the immobilization of glucose oxidase retaining its activity. Electron transport property of CGNC has been exploited to enhance the electron transport between the analyte and electrode. Electrochemical characteristics of the biosensor were studied by cyclic voltammetry and chronoamperometry. Under optimal conditions the biosensor exhibits a linear range from 1 M to 100 M with a limit of detection (lod) down to 100 nM. The sensor shows a low Michaelis-Menten constant value of 2.4 M indicates the high affinity of enzyme to the analyte points to the retained activity of enzyme after immobilization. The present glucose sensor with the high selectivity, sensitivity and stability is promising for practical clinical applications. © 2013 Elsevier Ltd. All rights reserved.
1. Introduction Sensitive and accurate determination of glucose concentration is extremely important in diagnosis and treatment of diabetes mellitus, clinical biochemistry, waste water treatment and food industry. Owing to the high sensitivity, reliability, simple instrumentation, low cost, exceptional compatibility and miniaturization, electrochemical techniques become important tools in the detection strategy. Studies were focusing on the development of amperometric glucose sensors based on the glucose oxidase catalyzed conversion of glucose to gluconolactone [1–4]. Though this method provides greater selectivity, direct immobilization of the enzyme on the electrode surface may lead to the loss of activity of the enzyme. Major drawback of most of the methods reported are the difficulty of electron transport between the active site of glucose oxidase and the electrode since the flavin adenine dinucleotide (FAD) moiety is deeply buried inside a protective protein shell. Nowadays enzyme free glucose sensors based on the direct electrocatalytic oxidation of glucose at an electrode surface receive keen research interest. A variety of metallic (Pt, Pd, Au, Ni and Cu) nanostructures [5–8], alloys (Pt-Ni, Pt-Au) [9,10] and metal oxides have been extensively used for the direct electrooxidation of glucose. Among which, nickel based nanomaterials
∗ Corresponding author. Tel.: +91 495 2286537; fax: +91 495 2287250. E-mail addresses: [email protected], [email protected] (N. Sandhyarani). 0013-4686/$ – see front matter © 2013 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.electacta.2013.07.010
exhibit remarkable activity due to the presence of Ni(OH)2 /NiOOH redox couple. Different strategies like electrodeposition of NiO to MWCNT [11], dispersion of nickel nanoparticles in grapheme [9] and dropping of NiO nanofibres onto graphene oxide modified glassy carbon electrode have been adopted for the modification of traditional electrodes. However, the instability of the catalytic film under applied potential and reduced electron transport from the redox couple to metal electrode surface remains as challenging issues. Chitosan is an attractive material in the construction of amperometric biosensors owing to the excellent biocompatibility and biodegradability [12]. Ability to form good films of chitosan makes it an excellent immobilizing platform for biomolecules in recent years [13]. Recently, we demonstrated the efficient electron transport property of chitosan gold nanocomposite (CGNC) film and their application in the development of a heavy metal sensor [14]. Utilizing the merits of glucose oxidase, Ni2+ /Ni3+ redox couple and CGNC film, herein we report a novel strategy for the development of a highly sensitive glucose sensor. The sensor was fabricated by immobilizing glucose oxidase to gold electrode modified with CGNC and Ni(OH)2 . This multilayered configuration integrated the functions of an enzyme immobilization matrix, suitable electrocatalyst and an efficient electron transport medium enhancing the sensitivity and selectivity of the biosensor. The application of such a system in the measurement of real sample also has been evaluated under optimal conditions. The exact functions of the enzyme and Ni(OH)2 in this integrated system are proposed.
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2. Experimental
2.4. Characterizations
2.1. Materials
Cyclic voltammetric (CV) analysis was performed with CH 400A Electrochemical Quartz Crystal Microbalance (CH Instruments, Austin, Texas). The three electrode system consisted of a modified gold electrode as working electrode, platinum wire as auxiliary electrode and saturated calomel electrode as the reference electrode against which all potentials were measured. Electrochemical measurements were carried out at room temperature with a scan rate of 100 mV s−1 . Energy dispersive X-ray spectroscopy (EDS) results was obtained using Hitachi SU6600 Variable Pressure Field Emission Scanning Electron Microscope. Attenuated total reflectance infrared spectra (ATR-IR) were recorded using a Perkin-Elmer spectrophotometer, model SPECTRUM TWO. Atomic force microscope (AFM) images and surface roughness of the electrodes were obtained using Park XE-100 AFM. The measurements were acquired in contact mode with silicon nitride cantilever. X-ray photo electron spectroscopic measurements were obtained using XPS, Axis Ultra, Shimadzu with Al K␣ monochromatic radiation.
Chloroauric acid (HAuCl4 ) and sodium borohydride were purchased from SISCO Laboratories, India and were used as received. Acetic acid, sodium hydroxide and nickel chloride were obtained from Merck International. Glucose oxidase from Aspergillus niger and Peroxidase from horse radish type V1-A were purchased from Sigma–Aldrich, India. Biomedical grade chitosan was obtained from the Central Institute of Fisheries Technology, Cochin, India. Chitosan has a molecular weight of 270 kDa, and degree of deacetylation was 85%. This was used as received. All aqueous solutions were prepared with ultra pure water obtained from an ultra filtration system (18 M cm. Milli-Q, Millipore system). 2.2. Synthesis of chitosan gold nanocomposite (CGNC) CGNC was synthesized using chloroauric acid and chitosan [14]. Briefly, 12 mL of a 1% (1 g/100 mL) chitosan solution was prepared in 2% (2 mL/100 mLwater) acetic acid. 25 mL of a 5 mM chloroauric acid solution was added to the vigorously stirred solution of chitosan. The resulting mixture was stirred for 1 h, and 2 mL of 10 mM ice cold sodium borohydride was added drop wise. Stirring was continued for another 2 h. The solution was centrifuged to obtain CGNC pellet, washed many times with acidic water and purified. 2.3. Preparation of electrodes Prior to the surface coating, the gold electrode was abraded with emery paper and polished with slurry of 0.3 m ␣-Al2 O3 powder and rinsed with ultra pure water. The electrode was immersed in a freshly prepared piranha solution for 10 min and sonicated with deionised water, rinsed with ethanol and dried. In this study, we explored the electron transport property of gold nanoparticles for electrochemical biosensing of glucose by depositing a thin layer of CGNC on gold electrode. A thin film of CGNC has been fabricated on the gold electrode surface by electrodeposition [15]. For electrodeposition two clean gold electrodes were dipped into the aqueous CGNC solution and connected to a DC power supply of 3 V. CGNC solution of weight percentage ∼0.5 mg/20 mL was used for the deposition. In order to see whether the concentration of CGNC has any important role on film thickness, we have used three different weight percentages (0.5 mg/20 mL, 0.25 mg/20 mL and 0.125 mg/20 mL) of CGNC for electrodeposition. We observed the same voltammogram for Fe2+ /Fe3+ redox couple in all the three electrodes indicated that the concentration of CGNC has no significant role in determining the film thickness, however the deposition time plays a very important role. Film thickness was controlled by varying the deposition time from 30 s to 5 min and obtained maximum electron transport with a deposition time of 1 min (Figure S1 of supplementary information). Hence we applied the potential for 1 min to attain an optimum thickness of CGNC on gold electrode. After deposition, the electrodes were removed, washed with distilled water and dried in a desiccator overnight. In the second step, electrodeposition of metallic nickel was carried out in 5 mM nickel chloride solution at a constant potential of 0.9 V for 200 s. Oxidation of metallic nickel to nickel hydroxide was achieved by cyclic voltammetry scanning from 0.2 V to 0.6 V in 0.1 M NaOH solution at a scan rate of 100 mV s−1 for 40 cycles [16]. Glucose oxidase was then successfully immobilized on the Ni(OH)2 thin film via physical adsorption. The prepared bio electrode was kept overnight for drying. The sensor will be represented as Au/CGNC/Ni(OH)2 /GOx .
3. Results and discussion 3.1. Characterization of electrode surface 3.1.1. Cyclic voltammetry Cyclic voltammetry was used for the electrochemical evaluation of the surface. Voltammetric response of Au/CGNC/Ni(OH)2 /GOx electrode has been monitored in 5 mM Fe(CN)6 3−/4− and 0.1 M KCl at 100 mV s−1 . Figure S2 of supplementary information shows the cyclic voltammogram of Au/CGNC/Ni(OH)2 /GOx electrode in each step of fabrication. It was seen from the voltammogram that bare gold electrode exhibited well defined redox peaks due to Fe2+ /Fe3+ redox couple. A reduction of the faradaic current on the CGNC modified electrode confirms the deposition. Slight decrease in the redox current indicates the formation of a non porous, slightly insulating layer on the electrode surface. Electrodeposition of nickel hydroxide on CGNC does not alter the current density significantly. A feeble faradaic current after the immobilization of glucose oxidase shows the formation of an insulating layer on the electrode surface. This can be easily understood because the enzyme acts as a barrier in diffusion on the surface of electrode. The well defined nature of the peaks confirms the presence of homogenous film formation on the surface in each step. To find the optimum conditions for the effective deposition of Ni(OH)2 the cyclic voltammetric scan number was varied from 20 to 50 during the electrochemical conversion of metallic nickel to nickel hydroxide. In cycles 20, 30, 40 and 50 voltammogram of Ni2+ /Ni3+ redox couple was noted (Figure S3 of supplementary information) and analyzed for the calculation of surface coverage of Ni(OH)2 . The surface active concentration of Ni(OH)2 on the modified electrode was estimated according to Faraday’s law: Ip =
n2 F 2 A v 4RT
where, Ip is the reduction current, n the number of electrons transferred, F Faraday’s constant (96,493 C mol−1 ), A the electrode area (cm2 ), the surface average concentration of electroactive Ni(OH)2 (mol/cm2 ), the scan rate (Vs−1 ), R the gas constant (8.314 Jmol−1 K−1 ) and T the temperature in Kelvin (298 K). The surface active concentration of Ni(OH)2 on four different voltammetric cycles viz. 20, 30, 40 and 50 were 1.42 × 10−9 mol/cm2 , 3.03 × 10−9 mol/cm2 , 3.92 × 10−9 mol/cm2 and 3.81 × 10−9 mol/cm2 respectively. Since we observed almost same coverage with 40 and 50 cycles, we selected 40 voltammetric cycles as the optimum condition for the deposition of Ni(OH)2 .
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Table 1 Percentage of elements on each stage of immobilization as indicated by EDS. Electrode
Au Au/CGNC Au/CGNC/Ni(OH)2 Au/CGNC/Ni(OH)2 /GOx
Weight percentage of elements (%) Au
C
N
O
Ni
100.00 77.47 71.38 67.03
– 16.53 14.56 17.15
– 2.8 1.1 3.14
– 3.2 11.62 12.16
– – 1.33 0.52
3.1.2. Energy dispersive X-ray spectroscopy To ascertain the composition of elements present on the electrode surface, energy dispersive X-ray spectroscopic analysis (EDS) was performed. Table 1 indicates the weight percentage of elements obtained on successive stages of electrode modification. The weight percentage of gold decreased from 100 to 67.03 as each layer added to the gold electrode. The observed changes in weight percentage of gold indicate the formation of different layers on the surface. Changes observed in the weight percentages of other elements further confirm the successful immobilization of various layers. AFM (supplementary information Figure S4a) and SEM (supplementary information Figure S4b) were used for the morphological characterization of the modified electrode surface. 3.1.3. ATR-IR spectroscopy ATR-IR was employed for the molecular characterization of surface (supplementary information Figure S5). The spectrum of Au/CGNC shows the characteristic peaks at 3310–3243 cm−1 of (NH2 ), 2922 cm−1 of (CH2 ) and 1077 cm−1 of (C O C). Upon the deposition of Ni(OH)2 the N H stretching peak of chitosan was completely masked by a broad peak at 3296 cm−1 . The broad band centred at 3296 cm−1 is owing to the O H vibration of hydrogen bonded water molecules located in the interlamellar space of Ni(OH)2 . The broad nature of the OH stretching indicates that the Ni(OH)2 formed on the electrode surface is in the ␣-form [17]. Intensity of other characteristic peaks of chitosan is decreased in the spectrum confirming the successful immobilization of ␣-Ni(OH)2 on the sensor surface. Further immobilization of glucose oxidase decreased the intensity of almost all the peaks and the characteristic peak of amide linkage in the enzyme was observed at 1657 cm−1 and 1547 cm−1 (Amide I and Amide II bands) confirming the presence of enzyme on the sensor surface. 3.1.4. X-ray photo electron spectroscopy The elemental characterization of the sensor surface was probed using X-ray photoelectron spectroscopy. Au 4f7/2 and 4f5/2 peaks are observed at 84.3 eV and 88 eV and C 1s, N 1s, O 1s binding energies were observed at 286 eV, 400 eV and 533 eV respectively. In Fig. 1 we show the survey scan and Ni 2P3/2 region showing the characteristic binding energy of Ni2+ . Major peak was observed at 857.4 eV with a satellite peak at 862.6 eV which is assigned as to the Ni(OH)2 . Two shoulder peaks at 855.7 eV and 859 eV were observed. The less intense peak at 855.7 eV is due to the presence of Ni in NiO [18]. The shoulder peak at 859 eV may be due to the presence of Ni(OH)2 on the surface. Note that the intensities of these shoulder peaks are very less. We have also characterized the sensor surface with XRD. Though most of the characteristic Ni(OH)2 peaks were masked by the presence of gold reflections a broad peak centred at 19◦ was present which again indicate the formation of ␣-Ni(OH)2 on the surface (not shown). It is known that when nickel hydroxide deposited using a potential up to 600 mV yields ␣-Ni(OH)2 [19]. However, ␣-Ni(OH)2 is not stable under alkaline solutions and slowly get converted to -Ni(OH)2 as a result there is a possibility of formation of small amount of -Ni(OH)2 during the measurement. Also upon
ageing the conversion from ␣-Ni(OH)2 to -Ni(OH)2 takes place [20]. In this work all the characterizations suggest the formation of ␣-Ni(OH)2 on the surface in the conditions we used for the deposition. Having understood the characteristics of the sensor surface we have analyzed the electrocatalytic reaction of glucose on the sensor surface. The experimental conditions such as pH of the experiments were optimized by performing the measurements at various pH values. Optimization of the pH of the solution is significant in this experiment as the performance of Ni(OH)2 be maximum at alkaline pH and the enzyme activity will be maximum at pH 7 [21,22]. We investigated the sensor response at pH values from 7.8 to 11.8 by varying the concentration of NaOH (Figure S6 of supplementary information). Prominent Ni2+ /Ni3+ redox peaks were observed only at pH 11.8 (in 0.1 M NaOH). The peaks were observed stable even up to 20 cycles. Hence we selected 0.1 M NaOH as the electrolyte for our further studies. In this study we are monitoring the peaks of Ni(OH)2 for determining the presence of glucose. Under the same conditions glucose, hydrogen peroxide or glucose oxidase did not give any observable voltammogram. 3.2. Electrocatalytic oxidation of glucose 3.2.1. Electrochemical response of the sensor in the absence of glucose oxidase The exact functions and importance of various layers in the integrated system have been systematically investigated through following sets of experiments. In order to study the need of a matrix for nickel hydroxide deposition, Ni(OH)2 has been directly deposited on gold electrode. Cyclic voltammogram indicates that the formation of Ni(OH)2 is not effective on bare gold electrode as no redox peaks were observed (supplementary information Figure S7). To ensure the uniform deposition of Ni(OH)2 , we selected chitosan as a matrix owing to its biocompatibility, excellent film forming capability along with the affinity to nickel [12]. The fabricated sensor is denoted as Au/Chit/Ni(OH)2 . Well defined redox peaks of Ni2+ /Ni3+ were observed in this electrode surface and the glucose detection was tested. Fig. 2A presents the CV response of Au/Chit/Ni(OH)2 electrode in 0.1 M NaOH containing glucose of different concentrations. The lowest detection limit of glucose was 100 M. In a previous study [14] we observed that the insulating property of chitosan film partially block the electron transport from the electrolyte to metal electrode; however a composite of chitosan gold nanoparticles (CGNC) facilitate the electron transport owing to the electron transport property of gold nanoparticles. Accordingly in this investigation we replaced chitosan film with CGNC film to improve the sensitivity of the sensor. The voltammetric response of the modified sensor Au/CGNC/Ni(OH)2 towards glucose is given in Fig. 2B. A 100 fold increase in the lower detection limit (1 M for Au/CGNC/Ni(OH)2 and 100 M for Au/Chit/Ni(OH)2 electrode) was observed while using Au/CGNC/Ni(OH)2 surface indicating the enhanced electron transport property of CGNC, which facilitate the electron transport between Ni2+ /Ni3+ redox couple to the gold electrode. In both the above cases in the absence of glucose, a couple of well defined Ni2+ /Ni3+ redox peaks were observed. Upon the addition of glucose, an increase in the anodic and cathodic peak current was noted. This is in consistent with other Ni(OH)2 based glucose sensors [23–25]. This electrochemical behaviour of Ni(OH)2 based sensors towards glucose is different from some other nickel based nonenzymatic glucose sensors where the cathodic peak current is expected to decrease in presence of glucose [26,27]. Based on the information about the sensor in this study we propose a detailed mechanism as explained below. It is reported that
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Fig. 1. X-ray photoelectron spectra of the sensor surface (A) survey scan and (B) Ni 2p3/2 region. The intense peak confirms the Ni(OH)2 on the surface.
on using a potential up to 600 mV for the measurement, the ␣Ni(OH)2 forms on the surface undergoes a reversible reaction to form ␥-NiO(OH) [19]. In our experiment we used a potential up to 600 mV and confirmed the presence of ␣-Ni(OH)2 on the surface. The reactions on the electrode surface are, ␣-Ni(OH)2 + OH− ↔ ␥-NiO(OH) + H2 O + e−
(1)
␥-NiO(OH) + glucose ↔ ␣-Ni(OH)2 + gluconolactone
(2)
␥-NiO(OH) reacts with glucose (2nd reaction) and regenerate the electro active Ni(OH)2 with predominant ␣-structure. The ␣Ni(OH)2 formed in the 2nd reaction also get converted to ␥-NiO(OH) along with the redox reaction on reacting with OH− from the electrolyte. This accelerates the oxidation reaction of the redox reaction 1 resulting in an increasing anodic current. This contributes to a higher rate of formation and concentration of Ni3+ on the surface. Thus formed Ni3+ converts back to Ni2+ in the reduction reaction and through oxidation of glucose which resulted in an increase of cathodic current. Thus the rates of reactions in the redox couple in presence of glucose will be higher than that in the absence of glucose which resulted in an increase of current in both oxidation and reduction peaks. Few of the ␥-NiO(OH) on reaction with glucose may also get converted to -Ni(OH)2 which in turn get converted to NiO or NiH in an irreversible reaction. This is possible as we perform the experiment in alkaline solution where ␣-Ni(OH)2 is not stable. However, since large number of Ni2+ and Ni3+ are present on the surface these smaller conversions may not affect the redox reaction and current significantly. It may be noted that the coverage of nickel hydroxide is very high on the surface and in the low concentration of glucose more than sufficient number of Ni(OH)2 and
NiO(OH) are available on the surface to react with. It is clear from the figure that as the glucose concentration increases the increase in anodic current becomes more compared to the increase in cathodic current. This is due to the consumption of more and more Ni3+ for the reaction with glucose resulted in the increase of oxidation peak current. The change may also be due to the formation of more -Ni(OH)2 . In order to evaluate the selectivity of the sensor, the interference effect of 1 mM hydrogen peroxide (H2 O2 ), 1 mM bovine serum albumin (BSA) and 1 mM ascorbic acid (AA) was studied. Current response of such electroactive interfering species is found to be negligible in comparison to that of glucose (data not shown). 3.2.2. Sensor performance in the presence of glucose oxidase To attain more specific and sensitive detection of glucose, we immobilized the enzyme glucose oxidase (GOx ) on to the best performing Au/CGNC/Ni(OH)2 electrode. The sensing response of Au/CGNC/Ni(OH)2 /GOx electrode as a function of glucose concentration is given in Fig. 3A. Here also we observed an increase for both anodic and cathodic current. By the use of enzyme on the surface, limit of detection was increased to 100 nM, which is one of the lowest reported detection limit for glucose. We are not comparing the absolute value of the redox current between the electrodes as the surface area and hence the current will be different for different electrodes. The sensing mechanism is a combination of glucose oxidase induced oxidation of glucose and Ni2+ /Ni3+ redox reaction (see below). Possibility of the interference of the oxidation of hydrogen peroxide at potential 0.5 V (which may form during the reaction of glucose with glucose oxidase) was ruled out by performing an experiment with Au/CGNC/GOx (without nickel hydroxide).
Fig. 2. Electrochemical response of (A) Au/Chit/Ni(OH)2 (B) Au/CGNC/Ni(OH)2 in 0.1 M NaOH containing glucose of varying concentration.
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Fig. 3. (A) Cyclic voltammogram of Au/CGNC/Ni(OH)2 /GOx sensor in glucose solution of different concentration (B) CV response of Au/CGNC/Ni(OH)2 /GOx sensor at scan rates 0.01,0.03, 0.05, 0.08, 0.1, 0.13, 0.15, 0.18 and 0.2 V/s respectively in 1 mM glucose solution.
We noted that no peaks were observed at this potential due to the oxidation of H2 O2 (Figure S8 of supplementary information). Fig. 3B shows the voltammogram of Au/CGNC/Ni(OH)2 /GOx electrode in 1 mM glucose solution at different scan rate. The peak to peak separation widened and the peak potentials increased with increasing scan rate. The linear increase of anodic and cathodic current indicates a surface controlled electrochemical reaction. From Figs. 2 and 3 it is clear that the peaks observed are due to the nickel redox reaction. We observed an enhancement in sensitivity due to the inclusion of glucose oxidase, which shows that glucose oxidase clearly play a major role in the detection. We propose the following mechanism for the reaction and detection of glucose in the integrated sensor which is different from the above said mechanism: GOx (FAD) + Glucose → GOx (FADH2 ) + Gluconolactone
(4)
2˛-Ni(OH)2 + 2OH− ↔ 2-NiO(OH) + 2H2 O + 2e−
(5)
GOx (FADH2 ) + 2-NiO(OH) → GOx (FAD) + 2˛-Ni(OH)2
(6)
The enzyme glucose oxidase oxidizes glucose to gluconolactone and the enzyme gets reduced in this process. The reduced species transferred the excess electrons to Ni2+ /Ni3+ redox couple leading to an enhancement in the redox current. It is known that one electron acceptors are good substrates for glucose oxidase [28] and the one electron acceptor and reduced form of enzyme exchange the protons and electrons. The explanation we offered in the previous section is valid here too as the nickel hydroxide is same in both the cases. The ␥-NiO(OH) reacts with GOx (FADH2 ) to form Ni(OH)2 (predominant form will be ␣ and few with ). Increase in current with increasing concentration of glucose is attributed to
the increased rate of reaction as explained above. Difference here is that the redox couple reacts with the GOx (FADH2 ) instead with glucose directly. The mechanism is indicated in Scheme 1. The stability of the deposited layers on the electrode surface has been examined by recording the current response of the redox couple in 1 mM glucose solution after 20 consecutive voltammetric cycles. The same current response has been observed even after a continuous scan of 20 cycles (Figure S9 of supplementary information) which demonstrates that the modified electrode is very stable. There is a possibility of GOx (FADH2 ) reacting with oxygen to form hydrogen peroxide at high concentration of glucose (supplementary information, Figure S10). It was also seen that the concentration of glucose oxidase on the surface is very important in determining the limit of detection of glucose (see supplementary information, Figure S11) which clearly indicate the role of GOx in the sensing of glucose.
3.3. Amperometric measurements and determination of Michaelis–Menten constant To determine the linear range of the proposed sensor, chronoamperometric analysis was employed at an applied potential of 0.4 V in the concentration range from 100 nM to 10 mM. It can be seen that the sensor exhibited a good linear response (R2 = 0.997) to glucose with a high sensitivity of 16,840 A mM−1 cm−2 from 1 M to 100 M (Fig. 4A). Michaelis–Menten constant value for Au/CGNC/Ni(OH)2 /GOx electrode calculated from the electrochemical version of Lineweaver–Burk plot (Fig. 4B) was 2.4 M. The low app KM value indicates the high affinity of the enzyme glucose oxidase to the analyte. The low value suggests that the enzyme retains its activity during the immobilization process which re affirm the role
Scheme 1. Representation of the mechanism of sensing of glucose on the integrated sensor.
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Fig. 4. (A) Linear calibration curve of Au/CGNC/Ni(OH)2 /GOx electrode for glucose detection (B) Lineweaver–Burk plot and (C) Selectivity of the sensor towards glucose.
of CGNC and Ni(OH)2 as a biomatrix for the enzyme immobilization. The selectivity of Au/CGNC/Ni(OH)2 /GOx electrode towards glucose has been studied by monitoring the CV response of common interferants viz. 1 mM hydrogen peroxide, 1 mM ascorbic acid and 1 mM bovine serum albumin (Fig. 4C). The result indicates that the sensor is highly selective for glucose. To evaluate the fabrication reproducibility of the sensor surface, four different electrode surfaces has been prepared and measured their current response towards the oxidation of 1 mM glucose (Figure S12 of supplementary information). The sensor surfaces showed a RSD of 1.39%. Storage stability of the sensor was evaluated by measuring the current response to 1 mM glucose. It was found that the sensor retained about 85% of its initial response after 20 days of storage. In order to demonstrate the utilization of proposed sensor in practical clinical application, it was applied for the determination of glucose concentration in human blood serum samples. Blood serum sample has been obtained from a health clinic. 10 L of blood serum sample was added to 10 mL of 0.1 M NaOH and the chronoamperometric response was recorded at 0.4 V to obtain the steady state current. Concentration of glucose corresponding to the steady state current has been evaluated from the standard calibration curve (Fig. 4A). The average glucose concentration determined was 84.96 mg dL−1 with a relative standard deviation of 0.64% which is in well agreement with the value measured from the local health clinic (84.7 mg dL−1 ). The performance of the sensor is compared with some of the existing sensors as shown in table S1 of supplementary information. It can be observed that the sensor presented in this paper show one of the lowest limits of detection and highest sensitivity.
4. Conclusions In summary, a combination of Ni2+ /Ni3+ redox couple and glucose oxidase has successfully been exploited for the realization of a highly sensitive glucose sensor for the first time. Chitosan gold nanoparticle composite film has been served as a bio matrix for the uniform film formation for Ni(OH)2 and a mediator for electron transport from the analyte to the surface. The combination of Ni(OH)2 and CGNC film acted as the immobilization platform for the enzyme in its active form. The proposed Au/CGNC/Ni(OH)2 /GOx sensor exhibited a lower detection limit of 100 nM with a high app sensitivity of 16,840 A mM−1 cm−2 . Low KM value of 2.4 M indicates the high affinity of immobilized glucose oxidase to glucose. The proposed sensor can be efficiently used in real practical clinical application.
Acknowledgements Financial support from Department of Science and Technology, India is gratefully acknowledged. M.M thanks CSIR, India for research fellowship. Mr. Priyadarshan and Dr. Gokulakrishnan Srinivasan of PerkinElmer are gratefully acknowledged for the ATRIR measurement.
Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.electacta. 2013.07.010.
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References [1] K. Wang, J.J. Xu, H.Y. Chen, A novel glucose biosensor based on the nanoscaled cobalt phthalocyanine–glucose oxidase biocomposite, Biosens. Bioelectron. 20 (2005) 1388. [2] M. Ammam, E.B. Easton, High-performance glucose sensor based on glucose oxidase encapsulated in new synthesized platinum nanoparticles supported on carbon Vulcan/Nafion composite deposited on glassy carbon, Sens. Actuators, B 155 (2011) 340. [3] P.P. Joshi, S.A. Merchant, Y. Wang, D.W. Schmidtke, Amperometric biosensors based on redox polymer-carbon nanotube-enzyme composites, Anal. Chem. 77 (2005) 3183. [4] X. Xiao, B. Zhou, L. Zhu, L. Xu, L. Tan, H. Tang, Y. Zhang, Q. Xie, S. Yao, An reagentless glucose biosensor based on direct electrochemistry of glucose oxidase immobilized on poly (methylene blue) doped silica nanocomposites, Sens. Actuators, B 165 (2012) 126. [5] E. Scavetta, S. Stipa, D. Tonelli, Electrodeposition of a nickel-based hydrotalcite on Pt nanoparticles for ethanol and glucose sensing, Electrochem. Commun. 9 (2007) 2838. [6] L.M. Lu, H.B. Li, F. Qu, X.B. Zhang, G.L. Shen, R.Q. Yu, In situ synthesis of palladium nanoparticle–graphene nanohybrids and their application in nonenzymatic glucose biosensors, Biosens. Bioelectron. 26 (2011) 3500. [7] Y. Liu, D. Yu, C. Zeng, Z. Miao, L. Dai, Biocompatible graphene oxide-based glucose biosensors, Langmuir 26 (2010) 6158. [8] H. Bai, M. Han, Y. Du, J. Bao, Z. Dai, Facile synthesis of porous tubular palladium nanostructures and their application in a nonenzymatic glucose sensor, Chem. Commun. 46 (2010) 1739. [9] H. Gao, F. Xiao, C.B. Ching, H. Duan, One-step electrochemical synthesis of PtNi nanoparticle-graphene nanocomposites for nonenzymatic amperometric glucose detection, ACS Appl. Mater. Interfaces 3 (2011) 3049. [10] J. Ryu, K. Kim, H.S. Kim, H.T. Hahn, D. Lashmore, Intense pulsed light induced platinum–gold alloy formation on carbon nanotubes for non-enzymatic glucose detection, Biosens. Bioelectron. 26 (2010) 602. [11] M. Shamsipur, M. Najafi, M.R.M. Hosseini, Highly improved electrooxidation of glucose at a nickel (II) oxide/multi-walled carbon nanotube modified glassy carbon electrode, Bioelectrochemistry 77 (2010) 120. [12] L. Luo, Q. Li, Y. Xu, Y. Ding, X. Wang, D. Deng, Y. Xu, Amperometric glucose biosensor based on NiFe2 O4 nanoparticles and chitosan, Sens. Actuators, B 145 (2010) 293. [13] Q. Sheng, K. Luo, J. Zheng, H. Zhang, Enzymatically induced formation of neodymium hexacyanoferrate nanoparticles on the glucose oxidase/chitosan modified glass carbon electrode for the detection of glucose, Biosens. Bioelectron. 24 (2008) 429.
[14] M. Mathew, S. Sureshkumar, N. Sandhyarani, Synthesis and characterization of gold–chitosan nanocomposite and application of resultant nanocomposite in sensors, Colloids Surf., B 93 (2012) 143. [15] X.L. Luo, J.J. Xu, Q. Zhang, G.J. Yang, H.Y. Chen, Electrochemically deposited chitosan hydrogel for horseradish peroxidase immobilization through gold nanoparticles self-assembly, Biosens. Bioelectron. 21 (2005) 190. [16] N. Qiao, J. Zheng, Nonenzymatic glucose sensor based on glassy carbon electrode modified with a nanocomposite composed of nickel hydroxide and grapheme, Microchim. Acta 177 (2012) 103. [17] J. Soler-Illia, M. Jobbagy, A.E. Regazzoni, M.A. Blesa, Synthesis of nickel hydroxide by homogeneous alkalinization. Precipitation mechanism, Chem. Mater. 11 (1999) 3140. [18] N. Hernandez, R. Moreno, A.J.S. Herencia, J.L.G. Fierro, Surface behavior of nickel powders in aqueous suspensions, J. Phys. Chem. B 109 (2005) 4470. [19] J.R.C. Salgado, M.H.S. Andrade, J.C.P. Silva, J. Tonholo, A voltammetric study of ␣-and -hydroxides over nickel alloys, Electrochim. Acta 47 (2002) 1997. [20] M.C. Bernard, P. Bernard, M. Keddam, S. Senyarich, H. Takenouti, Characterisation of new nickel hydroxides during the transformation of ␣-Ni(OH)2 to -Ni(OH)2 by ageing, Electrochim. Acta 41 (1996) 91. [21] H. Nie, Z. Yao, X. Zhou, Z. Yang, S. Huang, Nonenzymatic electrochemical detection of glucose using well-distributed nickel nanoparticles on straight multi-walled carbon nanotubes, Biosens. Bioelectron. 30 (2011) 28. [22] A.E.G. Cass, G. Davis, G.D. Francis, H.A.O. Hill, W.J. Aston, I.J. Higgins, E.V. Plotkin, L.D.L. Scott, A.P.F. Turner, Ferrocene-mediated enzyme electrode for amperometric determination of glucose, Anal. Chem. 56 (1984) 667. [23] C. Zhao, C. Shao, M. Li, K. Jiao, Flow-injection analysis of glucose without enzyme based on electrocatalytic oxidation of glucose at a nickel electrode, Talanta 71 (2007) 1769. [24] A. Safavi, N. Maleki, E. Farjami, Fabrication of a glucose sensor based on a novel nanocomposite electrode, Biosens. Bioelectron. 24 (2009) 1655. [25] X. Li, J. Liu, X. Ji, J. Jiang, R. Ding, Y. Hu, A. Hu, X. Huang, Ni/Al layered double hydroxide nanosheet film grown directly on Ti substrate and its application for a nonenzymatic glucose sensor, Sens. Actuators, B 147 (2010) 241. [26] L. Zheng, J. Zhang, J. Song, Ni (II)–quercetin complex modified multiwall carbon nanotube ionic liquid paste electrode and its electrocatalytic activity toward the oxidation of glucose, Electrochim. Acta 54 (2009) 4559. [27] Y. Mu, D. Jia, Y. He, Y. Miao, H.L. Wu, Nano nickel oxide modified non-enzymatic glucose sensors with enhanced sensitivity through an electrochemical process strategy at high potential, Biosens. Bioelectron. 26 (2011) 2948. [28] V. Leskovac, S. Trivic, G. Wohlfahrt, J. Kandrac, D. Pericin, Glucose oxidase from Aspergillus niger: the mechanism of action with molecular oxygen, quinones, and one-electron acceptors, Int. J. Bichem. Cell Biol. 37 (2005) 731.
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Electrochimica Acta journal homepage: www.elsevier.com/locate/electacta
A highly sensitive electrochemical glucose sensor structuring with nickel hydroxide and enzyme glucose oxidase Manjusha Mathew, N. Sandhyarani ∗ Nanoscience Research Laboratory, School of Nano Science and Technology, National Institute of Technology Calicut, Calicut, Kerala, India
a r t i c l e
i n f o
Article history: Received 11 May 2013 Received in revised form 27 June 2013 Accepted 4 July 2013 Available online xxx Keywords: Electrochemical sensor Nickel hydroxide Chitosan gold nanocomposite Cyclic voltammetry Glucose oxidase
a b s t r a c t A multilayered glucose biosensor with enhanced electron transport was fabricated via the sequential electrodeposition of chitosan gold nanocomposite (CGNC) and nickel hydroxide (Ni(OH)2 ) on a bare gold electrode and subsequent immobilization of glucose oxidase. A thin film of Ni(OH)2 deposited on CGNC modified gold electrode serves as an electrochemical redox probe as well as a matrix for the immobilization of glucose oxidase retaining its activity. Electron transport property of CGNC has been exploited to enhance the electron transport between the analyte and electrode. Electrochemical characteristics of the biosensor were studied by cyclic voltammetry and chronoamperometry. Under optimal conditions the biosensor exhibits a linear range from 1 M to 100 M with a limit of detection (lod) down to 100 nM. The sensor shows a low Michaelis-Menten constant value of 2.4 M indicates the high affinity of enzyme to the analyte points to the retained activity of enzyme after immobilization. The present glucose sensor with the high selectivity, sensitivity and stability is promising for practical clinical applications. © 2013 Elsevier Ltd. All rights reserved.
1. Introduction Sensitive and accurate determination of glucose concentration is extremely important in diagnosis and treatment of diabetes mellitus, clinical biochemistry, waste water treatment and food industry. Owing to the high sensitivity, reliability, simple instrumentation, low cost, exceptional compatibility and miniaturization, electrochemical techniques become important tools in the detection strategy. Studies were focusing on the development of amperometric glucose sensors based on the glucose oxidase catalyzed conversion of glucose to gluconolactone [1–4]. Though this method provides greater selectivity, direct immobilization of the enzyme on the electrode surface may lead to the loss of activity of the enzyme. Major drawback of most of the methods reported are the difficulty of electron transport between the active site of glucose oxidase and the electrode since the flavin adenine dinucleotide (FAD) moiety is deeply buried inside a protective protein shell. Nowadays enzyme free glucose sensors based on the direct electrocatalytic oxidation of glucose at an electrode surface receive keen research interest. A variety of metallic (Pt, Pd, Au, Ni and Cu) nanostructures [5–8], alloys (Pt-Ni, Pt-Au) [9,10] and metal oxides have been extensively used for the direct electrooxidation of glucose. Among which, nickel based nanomaterials
∗ Corresponding author. Tel.: +91 495 2286537; fax: +91 495 2287250. E-mail addresses: [email protected], [email protected] (N. Sandhyarani). 0013-4686/$ – see front matter © 2013 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.electacta.2013.07.010
exhibit remarkable activity due to the presence of Ni(OH)2 /NiOOH redox couple. Different strategies like electrodeposition of NiO to MWCNT [11], dispersion of nickel nanoparticles in grapheme [9] and dropping of NiO nanofibres onto graphene oxide modified glassy carbon electrode have been adopted for the modification of traditional electrodes. However, the instability of the catalytic film under applied potential and reduced electron transport from the redox couple to metal electrode surface remains as challenging issues. Chitosan is an attractive material in the construction of amperometric biosensors owing to the excellent biocompatibility and biodegradability [12]. Ability to form good films of chitosan makes it an excellent immobilizing platform for biomolecules in recent years [13]. Recently, we demonstrated the efficient electron transport property of chitosan gold nanocomposite (CGNC) film and their application in the development of a heavy metal sensor [14]. Utilizing the merits of glucose oxidase, Ni2+ /Ni3+ redox couple and CGNC film, herein we report a novel strategy for the development of a highly sensitive glucose sensor. The sensor was fabricated by immobilizing glucose oxidase to gold electrode modified with CGNC and Ni(OH)2 . This multilayered configuration integrated the functions of an enzyme immobilization matrix, suitable electrocatalyst and an efficient electron transport medium enhancing the sensitivity and selectivity of the biosensor. The application of such a system in the measurement of real sample also has been evaluated under optimal conditions. The exact functions of the enzyme and Ni(OH)2 in this integrated system are proposed.
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2. Experimental
2.4. Characterizations
2.1. Materials
Cyclic voltammetric (CV) analysis was performed with CH 400A Electrochemical Quartz Crystal Microbalance (CH Instruments, Austin, Texas). The three electrode system consisted of a modified gold electrode as working electrode, platinum wire as auxiliary electrode and saturated calomel electrode as the reference electrode against which all potentials were measured. Electrochemical measurements were carried out at room temperature with a scan rate of 100 mV s−1 . Energy dispersive X-ray spectroscopy (EDS) results was obtained using Hitachi SU6600 Variable Pressure Field Emission Scanning Electron Microscope. Attenuated total reflectance infrared spectra (ATR-IR) were recorded using a Perkin-Elmer spectrophotometer, model SPECTRUM TWO. Atomic force microscope (AFM) images and surface roughness of the electrodes were obtained using Park XE-100 AFM. The measurements were acquired in contact mode with silicon nitride cantilever. X-ray photo electron spectroscopic measurements were obtained using XPS, Axis Ultra, Shimadzu with Al K␣ monochromatic radiation.
Chloroauric acid (HAuCl4 ) and sodium borohydride were purchased from SISCO Laboratories, India and were used as received. Acetic acid, sodium hydroxide and nickel chloride were obtained from Merck International. Glucose oxidase from Aspergillus niger and Peroxidase from horse radish type V1-A were purchased from Sigma–Aldrich, India. Biomedical grade chitosan was obtained from the Central Institute of Fisheries Technology, Cochin, India. Chitosan has a molecular weight of 270 kDa, and degree of deacetylation was 85%. This was used as received. All aqueous solutions were prepared with ultra pure water obtained from an ultra filtration system (18 M cm. Milli-Q, Millipore system). 2.2. Synthesis of chitosan gold nanocomposite (CGNC) CGNC was synthesized using chloroauric acid and chitosan [14]. Briefly, 12 mL of a 1% (1 g/100 mL) chitosan solution was prepared in 2% (2 mL/100 mLwater) acetic acid. 25 mL of a 5 mM chloroauric acid solution was added to the vigorously stirred solution of chitosan. The resulting mixture was stirred for 1 h, and 2 mL of 10 mM ice cold sodium borohydride was added drop wise. Stirring was continued for another 2 h. The solution was centrifuged to obtain CGNC pellet, washed many times with acidic water and purified. 2.3. Preparation of electrodes Prior to the surface coating, the gold electrode was abraded with emery paper and polished with slurry of 0.3 m ␣-Al2 O3 powder and rinsed with ultra pure water. The electrode was immersed in a freshly prepared piranha solution for 10 min and sonicated with deionised water, rinsed with ethanol and dried. In this study, we explored the electron transport property of gold nanoparticles for electrochemical biosensing of glucose by depositing a thin layer of CGNC on gold electrode. A thin film of CGNC has been fabricated on the gold electrode surface by electrodeposition [15]. For electrodeposition two clean gold electrodes were dipped into the aqueous CGNC solution and connected to a DC power supply of 3 V. CGNC solution of weight percentage ∼0.5 mg/20 mL was used for the deposition. In order to see whether the concentration of CGNC has any important role on film thickness, we have used three different weight percentages (0.5 mg/20 mL, 0.25 mg/20 mL and 0.125 mg/20 mL) of CGNC for electrodeposition. We observed the same voltammogram for Fe2+ /Fe3+ redox couple in all the three electrodes indicated that the concentration of CGNC has no significant role in determining the film thickness, however the deposition time plays a very important role. Film thickness was controlled by varying the deposition time from 30 s to 5 min and obtained maximum electron transport with a deposition time of 1 min (Figure S1 of supplementary information). Hence we applied the potential for 1 min to attain an optimum thickness of CGNC on gold electrode. After deposition, the electrodes were removed, washed with distilled water and dried in a desiccator overnight. In the second step, electrodeposition of metallic nickel was carried out in 5 mM nickel chloride solution at a constant potential of 0.9 V for 200 s. Oxidation of metallic nickel to nickel hydroxide was achieved by cyclic voltammetry scanning from 0.2 V to 0.6 V in 0.1 M NaOH solution at a scan rate of 100 mV s−1 for 40 cycles [16]. Glucose oxidase was then successfully immobilized on the Ni(OH)2 thin film via physical adsorption. The prepared bio electrode was kept overnight for drying. The sensor will be represented as Au/CGNC/Ni(OH)2 /GOx .
3. Results and discussion 3.1. Characterization of electrode surface 3.1.1. Cyclic voltammetry Cyclic voltammetry was used for the electrochemical evaluation of the surface. Voltammetric response of Au/CGNC/Ni(OH)2 /GOx electrode has been monitored in 5 mM Fe(CN)6 3−/4− and 0.1 M KCl at 100 mV s−1 . Figure S2 of supplementary information shows the cyclic voltammogram of Au/CGNC/Ni(OH)2 /GOx electrode in each step of fabrication. It was seen from the voltammogram that bare gold electrode exhibited well defined redox peaks due to Fe2+ /Fe3+ redox couple. A reduction of the faradaic current on the CGNC modified electrode confirms the deposition. Slight decrease in the redox current indicates the formation of a non porous, slightly insulating layer on the electrode surface. Electrodeposition of nickel hydroxide on CGNC does not alter the current density significantly. A feeble faradaic current after the immobilization of glucose oxidase shows the formation of an insulating layer on the electrode surface. This can be easily understood because the enzyme acts as a barrier in diffusion on the surface of electrode. The well defined nature of the peaks confirms the presence of homogenous film formation on the surface in each step. To find the optimum conditions for the effective deposition of Ni(OH)2 the cyclic voltammetric scan number was varied from 20 to 50 during the electrochemical conversion of metallic nickel to nickel hydroxide. In cycles 20, 30, 40 and 50 voltammogram of Ni2+ /Ni3+ redox couple was noted (Figure S3 of supplementary information) and analyzed for the calculation of surface coverage of Ni(OH)2 . The surface active concentration of Ni(OH)2 on the modified electrode was estimated according to Faraday’s law: Ip =
n2 F 2 A v 4RT
where, Ip is the reduction current, n the number of electrons transferred, F Faraday’s constant (96,493 C mol−1 ), A the electrode area (cm2 ), the surface average concentration of electroactive Ni(OH)2 (mol/cm2 ), the scan rate (Vs−1 ), R the gas constant (8.314 Jmol−1 K−1 ) and T the temperature in Kelvin (298 K). The surface active concentration of Ni(OH)2 on four different voltammetric cycles viz. 20, 30, 40 and 50 were 1.42 × 10−9 mol/cm2 , 3.03 × 10−9 mol/cm2 , 3.92 × 10−9 mol/cm2 and 3.81 × 10−9 mol/cm2 respectively. Since we observed almost same coverage with 40 and 50 cycles, we selected 40 voltammetric cycles as the optimum condition for the deposition of Ni(OH)2 .
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Table 1 Percentage of elements on each stage of immobilization as indicated by EDS. Electrode
Au Au/CGNC Au/CGNC/Ni(OH)2 Au/CGNC/Ni(OH)2 /GOx
Weight percentage of elements (%) Au
C
N
O
Ni
100.00 77.47 71.38 67.03
– 16.53 14.56 17.15
– 2.8 1.1 3.14
– 3.2 11.62 12.16
– – 1.33 0.52
3.1.2. Energy dispersive X-ray spectroscopy To ascertain the composition of elements present on the electrode surface, energy dispersive X-ray spectroscopic analysis (EDS) was performed. Table 1 indicates the weight percentage of elements obtained on successive stages of electrode modification. The weight percentage of gold decreased from 100 to 67.03 as each layer added to the gold electrode. The observed changes in weight percentage of gold indicate the formation of different layers on the surface. Changes observed in the weight percentages of other elements further confirm the successful immobilization of various layers. AFM (supplementary information Figure S4a) and SEM (supplementary information Figure S4b) were used for the morphological characterization of the modified electrode surface. 3.1.3. ATR-IR spectroscopy ATR-IR was employed for the molecular characterization of surface (supplementary information Figure S5). The spectrum of Au/CGNC shows the characteristic peaks at 3310–3243 cm−1 of (NH2 ), 2922 cm−1 of (CH2 ) and 1077 cm−1 of (C O C). Upon the deposition of Ni(OH)2 the N H stretching peak of chitosan was completely masked by a broad peak at 3296 cm−1 . The broad band centred at 3296 cm−1 is owing to the O H vibration of hydrogen bonded water molecules located in the interlamellar space of Ni(OH)2 . The broad nature of the OH stretching indicates that the Ni(OH)2 formed on the electrode surface is in the ␣-form [17]. Intensity of other characteristic peaks of chitosan is decreased in the spectrum confirming the successful immobilization of ␣-Ni(OH)2 on the sensor surface. Further immobilization of glucose oxidase decreased the intensity of almost all the peaks and the characteristic peak of amide linkage in the enzyme was observed at 1657 cm−1 and 1547 cm−1 (Amide I and Amide II bands) confirming the presence of enzyme on the sensor surface. 3.1.4. X-ray photo electron spectroscopy The elemental characterization of the sensor surface was probed using X-ray photoelectron spectroscopy. Au 4f7/2 and 4f5/2 peaks are observed at 84.3 eV and 88 eV and C 1s, N 1s, O 1s binding energies were observed at 286 eV, 400 eV and 533 eV respectively. In Fig. 1 we show the survey scan and Ni 2P3/2 region showing the characteristic binding energy of Ni2+ . Major peak was observed at 857.4 eV with a satellite peak at 862.6 eV which is assigned as to the Ni(OH)2 . Two shoulder peaks at 855.7 eV and 859 eV were observed. The less intense peak at 855.7 eV is due to the presence of Ni in NiO [18]. The shoulder peak at 859 eV may be due to the presence of Ni(OH)2 on the surface. Note that the intensities of these shoulder peaks are very less. We have also characterized the sensor surface with XRD. Though most of the characteristic Ni(OH)2 peaks were masked by the presence of gold reflections a broad peak centred at 19◦ was present which again indicate the formation of ␣-Ni(OH)2 on the surface (not shown). It is known that when nickel hydroxide deposited using a potential up to 600 mV yields ␣-Ni(OH)2 [19]. However, ␣-Ni(OH)2 is not stable under alkaline solutions and slowly get converted to -Ni(OH)2 as a result there is a possibility of formation of small amount of -Ni(OH)2 during the measurement. Also upon
ageing the conversion from ␣-Ni(OH)2 to -Ni(OH)2 takes place [20]. In this work all the characterizations suggest the formation of ␣-Ni(OH)2 on the surface in the conditions we used for the deposition. Having understood the characteristics of the sensor surface we have analyzed the electrocatalytic reaction of glucose on the sensor surface. The experimental conditions such as pH of the experiments were optimized by performing the measurements at various pH values. Optimization of the pH of the solution is significant in this experiment as the performance of Ni(OH)2 be maximum at alkaline pH and the enzyme activity will be maximum at pH 7 [21,22]. We investigated the sensor response at pH values from 7.8 to 11.8 by varying the concentration of NaOH (Figure S6 of supplementary information). Prominent Ni2+ /Ni3+ redox peaks were observed only at pH 11.8 (in 0.1 M NaOH). The peaks were observed stable even up to 20 cycles. Hence we selected 0.1 M NaOH as the electrolyte for our further studies. In this study we are monitoring the peaks of Ni(OH)2 for determining the presence of glucose. Under the same conditions glucose, hydrogen peroxide or glucose oxidase did not give any observable voltammogram. 3.2. Electrocatalytic oxidation of glucose 3.2.1. Electrochemical response of the sensor in the absence of glucose oxidase The exact functions and importance of various layers in the integrated system have been systematically investigated through following sets of experiments. In order to study the need of a matrix for nickel hydroxide deposition, Ni(OH)2 has been directly deposited on gold electrode. Cyclic voltammogram indicates that the formation of Ni(OH)2 is not effective on bare gold electrode as no redox peaks were observed (supplementary information Figure S7). To ensure the uniform deposition of Ni(OH)2 , we selected chitosan as a matrix owing to its biocompatibility, excellent film forming capability along with the affinity to nickel [12]. The fabricated sensor is denoted as Au/Chit/Ni(OH)2 . Well defined redox peaks of Ni2+ /Ni3+ were observed in this electrode surface and the glucose detection was tested. Fig. 2A presents the CV response of Au/Chit/Ni(OH)2 electrode in 0.1 M NaOH containing glucose of different concentrations. The lowest detection limit of glucose was 100 M. In a previous study [14] we observed that the insulating property of chitosan film partially block the electron transport from the electrolyte to metal electrode; however a composite of chitosan gold nanoparticles (CGNC) facilitate the electron transport owing to the electron transport property of gold nanoparticles. Accordingly in this investigation we replaced chitosan film with CGNC film to improve the sensitivity of the sensor. The voltammetric response of the modified sensor Au/CGNC/Ni(OH)2 towards glucose is given in Fig. 2B. A 100 fold increase in the lower detection limit (1 M for Au/CGNC/Ni(OH)2 and 100 M for Au/Chit/Ni(OH)2 electrode) was observed while using Au/CGNC/Ni(OH)2 surface indicating the enhanced electron transport property of CGNC, which facilitate the electron transport between Ni2+ /Ni3+ redox couple to the gold electrode. In both the above cases in the absence of glucose, a couple of well defined Ni2+ /Ni3+ redox peaks were observed. Upon the addition of glucose, an increase in the anodic and cathodic peak current was noted. This is in consistent with other Ni(OH)2 based glucose sensors [23–25]. This electrochemical behaviour of Ni(OH)2 based sensors towards glucose is different from some other nickel based nonenzymatic glucose sensors where the cathodic peak current is expected to decrease in presence of glucose [26,27]. Based on the information about the sensor in this study we propose a detailed mechanism as explained below. It is reported that
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Fig. 1. X-ray photoelectron spectra of the sensor surface (A) survey scan and (B) Ni 2p3/2 region. The intense peak confirms the Ni(OH)2 on the surface.
on using a potential up to 600 mV for the measurement, the ␣Ni(OH)2 forms on the surface undergoes a reversible reaction to form ␥-NiO(OH) [19]. In our experiment we used a potential up to 600 mV and confirmed the presence of ␣-Ni(OH)2 on the surface. The reactions on the electrode surface are, ␣-Ni(OH)2 + OH− ↔ ␥-NiO(OH) + H2 O + e−
(1)
␥-NiO(OH) + glucose ↔ ␣-Ni(OH)2 + gluconolactone
(2)
␥-NiO(OH) reacts with glucose (2nd reaction) and regenerate the electro active Ni(OH)2 with predominant ␣-structure. The ␣Ni(OH)2 formed in the 2nd reaction also get converted to ␥-NiO(OH) along with the redox reaction on reacting with OH− from the electrolyte. This accelerates the oxidation reaction of the redox reaction 1 resulting in an increasing anodic current. This contributes to a higher rate of formation and concentration of Ni3+ on the surface. Thus formed Ni3+ converts back to Ni2+ in the reduction reaction and through oxidation of glucose which resulted in an increase of cathodic current. Thus the rates of reactions in the redox couple in presence of glucose will be higher than that in the absence of glucose which resulted in an increase of current in both oxidation and reduction peaks. Few of the ␥-NiO(OH) on reaction with glucose may also get converted to -Ni(OH)2 which in turn get converted to NiO or NiH in an irreversible reaction. This is possible as we perform the experiment in alkaline solution where ␣-Ni(OH)2 is not stable. However, since large number of Ni2+ and Ni3+ are present on the surface these smaller conversions may not affect the redox reaction and current significantly. It may be noted that the coverage of nickel hydroxide is very high on the surface and in the low concentration of glucose more than sufficient number of Ni(OH)2 and
NiO(OH) are available on the surface to react with. It is clear from the figure that as the glucose concentration increases the increase in anodic current becomes more compared to the increase in cathodic current. This is due to the consumption of more and more Ni3+ for the reaction with glucose resulted in the increase of oxidation peak current. The change may also be due to the formation of more -Ni(OH)2 . In order to evaluate the selectivity of the sensor, the interference effect of 1 mM hydrogen peroxide (H2 O2 ), 1 mM bovine serum albumin (BSA) and 1 mM ascorbic acid (AA) was studied. Current response of such electroactive interfering species is found to be negligible in comparison to that of glucose (data not shown). 3.2.2. Sensor performance in the presence of glucose oxidase To attain more specific and sensitive detection of glucose, we immobilized the enzyme glucose oxidase (GOx ) on to the best performing Au/CGNC/Ni(OH)2 electrode. The sensing response of Au/CGNC/Ni(OH)2 /GOx electrode as a function of glucose concentration is given in Fig. 3A. Here also we observed an increase for both anodic and cathodic current. By the use of enzyme on the surface, limit of detection was increased to 100 nM, which is one of the lowest reported detection limit for glucose. We are not comparing the absolute value of the redox current between the electrodes as the surface area and hence the current will be different for different electrodes. The sensing mechanism is a combination of glucose oxidase induced oxidation of glucose and Ni2+ /Ni3+ redox reaction (see below). Possibility of the interference of the oxidation of hydrogen peroxide at potential 0.5 V (which may form during the reaction of glucose with glucose oxidase) was ruled out by performing an experiment with Au/CGNC/GOx (without nickel hydroxide).
Fig. 2. Electrochemical response of (A) Au/Chit/Ni(OH)2 (B) Au/CGNC/Ni(OH)2 in 0.1 M NaOH containing glucose of varying concentration.
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Fig. 3. (A) Cyclic voltammogram of Au/CGNC/Ni(OH)2 /GOx sensor in glucose solution of different concentration (B) CV response of Au/CGNC/Ni(OH)2 /GOx sensor at scan rates 0.01,0.03, 0.05, 0.08, 0.1, 0.13, 0.15, 0.18 and 0.2 V/s respectively in 1 mM glucose solution.
We noted that no peaks were observed at this potential due to the oxidation of H2 O2 (Figure S8 of supplementary information). Fig. 3B shows the voltammogram of Au/CGNC/Ni(OH)2 /GOx electrode in 1 mM glucose solution at different scan rate. The peak to peak separation widened and the peak potentials increased with increasing scan rate. The linear increase of anodic and cathodic current indicates a surface controlled electrochemical reaction. From Figs. 2 and 3 it is clear that the peaks observed are due to the nickel redox reaction. We observed an enhancement in sensitivity due to the inclusion of glucose oxidase, which shows that glucose oxidase clearly play a major role in the detection. We propose the following mechanism for the reaction and detection of glucose in the integrated sensor which is different from the above said mechanism: GOx (FAD) + Glucose → GOx (FADH2 ) + Gluconolactone
(4)
2˛-Ni(OH)2 + 2OH− ↔ 2-NiO(OH) + 2H2 O + 2e−
(5)
GOx (FADH2 ) + 2-NiO(OH) → GOx (FAD) + 2˛-Ni(OH)2
(6)
The enzyme glucose oxidase oxidizes glucose to gluconolactone and the enzyme gets reduced in this process. The reduced species transferred the excess electrons to Ni2+ /Ni3+ redox couple leading to an enhancement in the redox current. It is known that one electron acceptors are good substrates for glucose oxidase [28] and the one electron acceptor and reduced form of enzyme exchange the protons and electrons. The explanation we offered in the previous section is valid here too as the nickel hydroxide is same in both the cases. The ␥-NiO(OH) reacts with GOx (FADH2 ) to form Ni(OH)2 (predominant form will be ␣ and few with ). Increase in current with increasing concentration of glucose is attributed to
the increased rate of reaction as explained above. Difference here is that the redox couple reacts with the GOx (FADH2 ) instead with glucose directly. The mechanism is indicated in Scheme 1. The stability of the deposited layers on the electrode surface has been examined by recording the current response of the redox couple in 1 mM glucose solution after 20 consecutive voltammetric cycles. The same current response has been observed even after a continuous scan of 20 cycles (Figure S9 of supplementary information) which demonstrates that the modified electrode is very stable. There is a possibility of GOx (FADH2 ) reacting with oxygen to form hydrogen peroxide at high concentration of glucose (supplementary information, Figure S10). It was also seen that the concentration of glucose oxidase on the surface is very important in determining the limit of detection of glucose (see supplementary information, Figure S11) which clearly indicate the role of GOx in the sensing of glucose.
3.3. Amperometric measurements and determination of Michaelis–Menten constant To determine the linear range of the proposed sensor, chronoamperometric analysis was employed at an applied potential of 0.4 V in the concentration range from 100 nM to 10 mM. It can be seen that the sensor exhibited a good linear response (R2 = 0.997) to glucose with a high sensitivity of 16,840 A mM−1 cm−2 from 1 M to 100 M (Fig. 4A). Michaelis–Menten constant value for Au/CGNC/Ni(OH)2 /GOx electrode calculated from the electrochemical version of Lineweaver–Burk plot (Fig. 4B) was 2.4 M. The low app KM value indicates the high affinity of the enzyme glucose oxidase to the analyte. The low value suggests that the enzyme retains its activity during the immobilization process which re affirm the role
Scheme 1. Representation of the mechanism of sensing of glucose on the integrated sensor.
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Fig. 4. (A) Linear calibration curve of Au/CGNC/Ni(OH)2 /GOx electrode for glucose detection (B) Lineweaver–Burk plot and (C) Selectivity of the sensor towards glucose.
of CGNC and Ni(OH)2 as a biomatrix for the enzyme immobilization. The selectivity of Au/CGNC/Ni(OH)2 /GOx electrode towards glucose has been studied by monitoring the CV response of common interferants viz. 1 mM hydrogen peroxide, 1 mM ascorbic acid and 1 mM bovine serum albumin (Fig. 4C). The result indicates that the sensor is highly selective for glucose. To evaluate the fabrication reproducibility of the sensor surface, four different electrode surfaces has been prepared and measured their current response towards the oxidation of 1 mM glucose (Figure S12 of supplementary information). The sensor surfaces showed a RSD of 1.39%. Storage stability of the sensor was evaluated by measuring the current response to 1 mM glucose. It was found that the sensor retained about 85% of its initial response after 20 days of storage. In order to demonstrate the utilization of proposed sensor in practical clinical application, it was applied for the determination of glucose concentration in human blood serum samples. Blood serum sample has been obtained from a health clinic. 10 L of blood serum sample was added to 10 mL of 0.1 M NaOH and the chronoamperometric response was recorded at 0.4 V to obtain the steady state current. Concentration of glucose corresponding to the steady state current has been evaluated from the standard calibration curve (Fig. 4A). The average glucose concentration determined was 84.96 mg dL−1 with a relative standard deviation of 0.64% which is in well agreement with the value measured from the local health clinic (84.7 mg dL−1 ). The performance of the sensor is compared with some of the existing sensors as shown in table S1 of supplementary information. It can be observed that the sensor presented in this paper show one of the lowest limits of detection and highest sensitivity.
4. Conclusions In summary, a combination of Ni2+ /Ni3+ redox couple and glucose oxidase has successfully been exploited for the realization of a highly sensitive glucose sensor for the first time. Chitosan gold nanoparticle composite film has been served as a bio matrix for the uniform film formation for Ni(OH)2 and a mediator for electron transport from the analyte to the surface. The combination of Ni(OH)2 and CGNC film acted as the immobilization platform for the enzyme in its active form. The proposed Au/CGNC/Ni(OH)2 /GOx sensor exhibited a lower detection limit of 100 nM with a high app sensitivity of 16,840 A mM−1 cm−2 . Low KM value of 2.4 M indicates the high affinity of immobilized glucose oxidase to glucose. The proposed sensor can be efficiently used in real practical clinical application.
Acknowledgements Financial support from Department of Science and Technology, India is gratefully acknowledged. M.M thanks CSIR, India for research fellowship. Mr. Priyadarshan and Dr. Gokulakrishnan Srinivasan of PerkinElmer are gratefully acknowledged for the ATRIR measurement.
Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.electacta. 2013.07.010.
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References [1] K. Wang, J.J. Xu, H.Y. Chen, A novel glucose biosensor based on the nanoscaled cobalt phthalocyanine–glucose oxidase biocomposite, Biosens. Bioelectron. 20 (2005) 1388. [2] M. Ammam, E.B. Easton, High-performance glucose sensor based on glucose oxidase encapsulated in new synthesized platinum nanoparticles supported on carbon Vulcan/Nafion composite deposited on glassy carbon, Sens. Actuators, B 155 (2011) 340. [3] P.P. Joshi, S.A. Merchant, Y. Wang, D.W. Schmidtke, Amperometric biosensors based on redox polymer-carbon nanotube-enzyme composites, Anal. Chem. 77 (2005) 3183. [4] X. Xiao, B. Zhou, L. Zhu, L. Xu, L. Tan, H. Tang, Y. Zhang, Q. Xie, S. Yao, An reagentless glucose biosensor based on direct electrochemistry of glucose oxidase immobilized on poly (methylene blue) doped silica nanocomposites, Sens. Actuators, B 165 (2012) 126. [5] E. Scavetta, S. Stipa, D. Tonelli, Electrodeposition of a nickel-based hydrotalcite on Pt nanoparticles for ethanol and glucose sensing, Electrochem. Commun. 9 (2007) 2838. [6] L.M. Lu, H.B. Li, F. Qu, X.B. Zhang, G.L. Shen, R.Q. Yu, In situ synthesis of palladium nanoparticle–graphene nanohybrids and their application in nonenzymatic glucose biosensors, Biosens. Bioelectron. 26 (2011) 3500. [7] Y. Liu, D. Yu, C. Zeng, Z. Miao, L. Dai, Biocompatible graphene oxide-based glucose biosensors, Langmuir 26 (2010) 6158. [8] H. Bai, M. Han, Y. Du, J. Bao, Z. Dai, Facile synthesis of porous tubular palladium nanostructures and their application in a nonenzymatic glucose sensor, Chem. Commun. 46 (2010) 1739. [9] H. Gao, F. Xiao, C.B. Ching, H. Duan, One-step electrochemical synthesis of PtNi nanoparticle-graphene nanocomposites for nonenzymatic amperometric glucose detection, ACS Appl. Mater. Interfaces 3 (2011) 3049. [10] J. Ryu, K. Kim, H.S. Kim, H.T. Hahn, D. Lashmore, Intense pulsed light induced platinum–gold alloy formation on carbon nanotubes for non-enzymatic glucose detection, Biosens. Bioelectron. 26 (2010) 602. [11] M. Shamsipur, M. Najafi, M.R.M. Hosseini, Highly improved electrooxidation of glucose at a nickel (II) oxide/multi-walled carbon nanotube modified glassy carbon electrode, Bioelectrochemistry 77 (2010) 120. [12] L. Luo, Q. Li, Y. Xu, Y. Ding, X. Wang, D. Deng, Y. Xu, Amperometric glucose biosensor based on NiFe2 O4 nanoparticles and chitosan, Sens. Actuators, B 145 (2010) 293. [13] Q. Sheng, K. Luo, J. Zheng, H. Zhang, Enzymatically induced formation of neodymium hexacyanoferrate nanoparticles on the glucose oxidase/chitosan modified glass carbon electrode for the detection of glucose, Biosens. Bioelectron. 24 (2008) 429.
[14] M. Mathew, S. Sureshkumar, N. Sandhyarani, Synthesis and characterization of gold–chitosan nanocomposite and application of resultant nanocomposite in sensors, Colloids Surf., B 93 (2012) 143. [15] X.L. Luo, J.J. Xu, Q. Zhang, G.J. Yang, H.Y. Chen, Electrochemically deposited chitosan hydrogel for horseradish peroxidase immobilization through gold nanoparticles self-assembly, Biosens. Bioelectron. 21 (2005) 190. [16] N. Qiao, J. Zheng, Nonenzymatic glucose sensor based on glassy carbon electrode modified with a nanocomposite composed of nickel hydroxide and grapheme, Microchim. Acta 177 (2012) 103. [17] J. Soler-Illia, M. Jobbagy, A.E. Regazzoni, M.A. Blesa, Synthesis of nickel hydroxide by homogeneous alkalinization. Precipitation mechanism, Chem. Mater. 11 (1999) 3140. [18] N. Hernandez, R. Moreno, A.J.S. Herencia, J.L.G. Fierro, Surface behavior of nickel powders in aqueous suspensions, J. Phys. Chem. B 109 (2005) 4470. [19] J.R.C. Salgado, M.H.S. Andrade, J.C.P. Silva, J. Tonholo, A voltammetric study of ␣-and -hydroxides over nickel alloys, Electrochim. Acta 47 (2002) 1997. [20] M.C. Bernard, P. Bernard, M. Keddam, S. Senyarich, H. Takenouti, Characterisation of new nickel hydroxides during the transformation of ␣-Ni(OH)2 to -Ni(OH)2 by ageing, Electrochim. Acta 41 (1996) 91. [21] H. Nie, Z. Yao, X. Zhou, Z. Yang, S. Huang, Nonenzymatic electrochemical detection of glucose using well-distributed nickel nanoparticles on straight multi-walled carbon nanotubes, Biosens. Bioelectron. 30 (2011) 28. [22] A.E.G. Cass, G. Davis, G.D. Francis, H.A.O. Hill, W.J. Aston, I.J. Higgins, E.V. Plotkin, L.D.L. Scott, A.P.F. Turner, Ferrocene-mediated enzyme electrode for amperometric determination of glucose, Anal. Chem. 56 (1984) 667. [23] C. Zhao, C. Shao, M. Li, K. Jiao, Flow-injection analysis of glucose without enzyme based on electrocatalytic oxidation of glucose at a nickel electrode, Talanta 71 (2007) 1769. [24] A. Safavi, N. Maleki, E. Farjami, Fabrication of a glucose sensor based on a novel nanocomposite electrode, Biosens. Bioelectron. 24 (2009) 1655. [25] X. Li, J. Liu, X. Ji, J. Jiang, R. Ding, Y. Hu, A. Hu, X. Huang, Ni/Al layered double hydroxide nanosheet film grown directly on Ti substrate and its application for a nonenzymatic glucose sensor, Sens. Actuators, B 147 (2010) 241. [26] L. Zheng, J. Zhang, J. Song, Ni (II)–quercetin complex modified multiwall carbon nanotube ionic liquid paste electrode and its electrocatalytic activity toward the oxidation of glucose, Electrochim. Acta 54 (2009) 4559. [27] Y. Mu, D. Jia, Y. He, Y. Miao, H.L. Wu, Nano nickel oxide modified non-enzymatic glucose sensors with enhanced sensitivity through an electrochemical process strategy at high potential, Biosens. Bioelectron. 26 (2011) 2948. [28] V. Leskovac, S. Trivic, G. Wohlfahrt, J. Kandrac, D. Pericin, Glucose oxidase from Aspergillus niger: the mechanism of action with molecular oxygen, quinones, and one-electron acceptors, Int. J. Bichem. Cell Biol. 37 (2005) 731.