Materials Science and Engineering C 40 (2014) 148–156
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
A novel functional conducting polymer as an immobilization platform Emine Guler a, Hakan Can Soyleyici b, Dilek Odaci Demirkol a,⁎, Metin Ak c, Suna Timur a,⁎ a b c
Ege University, Faculty of Science, Biochemistry Department, 35100 Bornova, Izmir, Turkey Adnan Menderes University, Faculty of Art and Science, Chemistry Department, Aydin, Turkey Pamukkale University, Faculty of Art and Science, Chemistry Department, Denizli, Turkey
a r t i c l e
i n f o
Article history: Received 15 December 2013 Received in revised form 18 February 2014 Accepted 22 March 2014 Available online 31 March 2014 Keywords: Conducting polymers Pyranose oxidase G. oxydans Biosensing
a b s t r a c t Here, we present the fabrication of conducting polymer based enzymatic and microbial biosensors. To obtain immobilization platforms for both pyranose oxidase (PyOx) and Gluconobacter oxydans, the graphite electrode surface was modified with the polymer of 4-amino-N-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzamide (HKCN) which has free amino groups on the surface for further bioconjugation reactions with the biomolecules. Initially, the electrode surface was covered with HKCN via electropolymerization. Then, either PyOx or G. oxydans cell was stabilized using glutaraldehyde as a cross-linker. After optimization of biosensors, analytical characterization and surface imaging studies were investigated. The change of current depends on glucose concentration between 0.05–1.0 mM and 0.25–2.5 mM with HKCN/PyOx and HKCN/G. oxydans biosensors in batch systems. Also, the calibration graphs were obtained for glucose in FIA mode, and in this case, linear ranges were found to be 0.01–1.0 mM and 0.1–7.5 mM for HKCN/PyOx and HKCN/G. oxydans, respectively. © 2014 Elsevier B.V. All rights reserved.
1. Introduction Biomolecule-based detection technology has become a dynamic area of research because of its remarkable potential for a variety of applications such as in food/beverage industries [1–4], diagnosis of disease [5–9], bioprocess monitoring [10–12] and screening of environmental pollutants [13–17]. To improve the stability of the biosensors for different requirements, modification of transducer surface for immobilization of bio-molecule is a key point. Different strategies such as self-assembly monolayers [18,19], natural polymers (gelatin, chitosan etc.) [20,21], sol–gel method [22], and conducting polymers [23–25] have been used to prepare immobilization platforms. The use of conducting polymers in the design of biosensors is very common because of homogeneous and manageable film character, ability of modification of physical and optical properties, stability and biocompatibility, availability of various types of monomers, reproducibility and easy production, and efficient electron transfer ability [26–28]. The synthesis of novel conducting polymers gives the creativity of the fabricated bio-detection systems. And also the addition of functional groups to the monomer backbone provides special and targeted immobilization of bio-components and also improves the stability of bio-components onto conducting polymers because of linkage via covalent attachment [29]. Here we described the use of a novel monomer with amino groups as an immobilization matrix. In the first part, 2,5-di(2-thienyl)pyrrole ⁎ Corresponding authors. Tel.: +90 2323115487; fax: +90 2323115485. E-mail addresses: [email protected], [email protected] (D.O. Demirkol), [email protected], [email protected] (S. Timur).
http://dx.doi.org/10.1016/j.msec.2014.03.063 0928-4931/© 2014 Elsevier B.V. All rights reserved.
derivative namely 4-amino-N-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl) benzamide (HKCN) has been synthesized via the reaction of 1,4-di(2thienyl)-1,4-butanedione and p-aminobenzoyl hydrazide. Using hydrazide instead of amine has not only increased product yield, but also improved properties of the corresponding polymer. Recently, optical properties of HKCN were described [30]. Herein, electrochemically deposited HKCN onto the graphite surfaces was used for the immobilization of PyOx and whole cell of Gluconobacter oxydans to fabricate electrochemical enzyme and whole cell biosensors. After optimization, the characterization studies such as linearity were carried out using both batch and FIA modes.
2. Materials and method 2.1. Reagents Pyranose oxidase (PyOx; pyranose:oxygen 2-oxidoreductase, E.C. 1.1.3.10, from Coriolus sp.), D-glucose, D(+)-xylose, ethanol, glycerol, and ascorbic acid were purchased from Sigma Chem. Co. and tetrabutylammoniumhexafluorophosphate (TBAPF6) was purchased from Aldrich. Dichloromethane (DCM) was obtained from Merck. Glutaraldehyde solution (25%, v/v) was purchased from Sigma-Aldrich. Commercial enzyme assay kit (Glucose MR, Cat. No. 1129010) was obtained from Cromatest. Aluminum chloride (AlCl3) (Aldrich), succinyl chloride (Aldrich), hydrochloric acid (Merck), NaHCO3 (Aldrich), MgSO4 (Aldrich), propionic acid (Aldrich), and toluene (Aldrich) were used for the synthesis of the monomer. All other chemicals were of analytical grade.
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2.2. Apparatus All amperometric measurements were carried out by PalmSens Electrochemical Measurement Unit (Palm Instruments, Houten, Netherlands). During batch operations, experiments were performed in a reaction cell (10 mL) at room temperature using a three electrode configuration, consisting of a graphite working electrode, a Ag/AgCl reference electrode (3.0 M KCl, Metrohm, Switzerland) and a platinum counter electrode (Metrohm, Switzerland). For the flow injection measurements the electrodes were mounted into an electrochemical flow-through cell of the cross-flow type with glassy carbon working, Ag/AgCl reference, and platinum wire counter electrodes (CHI130, Austin, www.chinstruments.com). The FIA system contains a peristaltic pump (FIAtron, Oconomowoc, WI, USA), an eightport injection valve (FIAtron, Oconomowoc, WI) and a cross-flow cell with three electrodes. The peristaltic pump equipped with Teflon tubing (0.50 mm inner diameter) carried 1.6 mL buffer solution per minute. Samples were injected with the eight-port injection valve with a 100 μL sample injection loop. The FIA system was connected to a PalmSens potentiostat for the electrochemical measurements. Colorimetric experiments were performed with a Pharmacia LKB Novaspec II spectrophotometer (LKB Biochrom, England). An Olympus BX53F fluorescence microscope, an Olympus DP72 camera and Uplanapo 100× objective were used in fluorescence images of monomer and prepared biosensors. 2.3. Cell cultivation The strain of G. oxydans DSMZ 2343 was obtained from the German Collection of Microorganisms and Cell Cultures (Braunschweig, Germany) and was sub-cultured on the agar containing (g/L): D-glucose, 100; yeast extract, 10; calcium carbonate, 20; agar, 20. Then, the cells were inoculated into liquid medium of the following composition, g/L: D-glucose, 5.0 and yeast extract, 5.0. Cells were grown at 28 °C on a rotary shaker (175 rpm) in 250 mL flasks with 50 mL medium. Cell growth was followed spectrophotometrically by measuring the optical density at 600 nm. Cells were collected in their late exponential phase, separated from culture medium by centrifugation (4000 rpm, 10 min), washed with sterile 0.9% NaCl solution and re-centrifuged. The obtained cellular paste was used for biosensor preparation [31]. 2.4. Synthesis of HKCN The monomer and its reagents (1,4-di(2-thienyl)-1,4-butanedione and p-aminobenzoyl hydrazide) were prepared according to the
Fig. 1. Repeated potential-scan electropolymerization of monomer in dichloromethane/ TBAPF6 (0.1 M) solvent–electrolyte system at a scan rate of 0.1 V/s on graphite (up to 10 cycles).
methods described in the literature [30]. All experiments were carried out under dry argon by using Standard Schlenk techniques. Solvents were dried, distilled and saturated with argon. A round-bottomed flask equipped with an argon inlet and magnetic stirrer was charged with 2.5 g (10 mmol) 1,4-di(2-thienyl)-1,4butanedione, 1.51 g (10 mmol) p-aminobenzoyl hydrazide, 0.2 g (1.2 mmol) PTSA, 1.0 mL DMSO and 20 mL toluene. The resultant mixture was stirred and refluxed for 18 h under argon. The darkened solution was filtered hot to remove oily decomposition products. Then, the mixture was cooled to room temperature and the solid product was filtered off to give a yellow powder that was washed with pentane (3 × 15 mL) and air-dried (yield: 3.1 g, 85%).
2.5. Preparation of HKCN/PyOx and HKCN/G. oxydans biosensors Before each experiment, graphite rods (Ringsdorff Werke GmbH, Bonn, Germany, 3.05 mm diameter and 13% porosity) were polished with wet emery paper and rinsed thoroughly with distilled water to be used in batch operations. Glassy carbon electrodes (GCEs) were cleaned by polishing with 0.05 μm alumina slurry followed by ultrasonication in ethanol and distilled water for 5 min prior to being used in flow injection mode of analysis (FIA).
Table 1 Comparison of analytical performance of conducting polymer based biosensors. Electrode
Conducting polymer
Biological material
Target analyte
Linearity
RSD
Reference
Graphite Graphite Graphite Graphite GCE Graphite Graphite Graphite Graphite Graphite Graphite
PBDT PESeE BEDOA-6 Poly(TBT6-NH2) PANI–CNT Poly(SNS-NH2) BlmTh:Fmoc-Gly-OH SNS-COOH SNS-COOH/Lys SNS-COOH/PAMAM G2 SNS-COOH/PAMAM G4
GOx GOx GOx ChO HRP AChE and ChO ChOx GOx GOx GOx GOx
Glucose Glucose Glucose Choline H2O2 Acetylcholine Cholesterol Glucose Glucose Glucose Glucose
0.05–2.0 mM 0.01–2.0 mM 0.025–1.25 mM 0.1–10.0 mM 0.2–19.0 μM 0.12–10.0 mM 0.3–10 μM 0.01–1.2 mM 0.01–2.40 mM 0.02–1.20 mM 0.02–1.20 mM
– – 5.87% 3.7% 4.6% 3.2% 4.4% – – – –
[35] [35] [36] [34] [37] [38] [39] [40] [28] [28] [28]
GCE: glassy carbon electrode, PBDT: poly(4,7-di(2,3)-dihydrothienol[3,4-b][1,4]dioxin-5-yl-benzo[1,2,5]thiadiazole), PESeE: poly(4,7-di(2,3)-dihydrothienol[3,4-b][1,4]dioxin-5-yl-2,1,3benzoselenadiazole), BEDOA-6: 6-(4,7-bis(2,3-dihydrothieno[3,4-b][1,4]dioxin-5-yl)-2H-benzo[d][1–3]triazol-2-yl)hexan-1-amine, poly(TBT6-NH2); poly(6-(4,7-di(thiophen-2-yl)2Hbenzo[d][1–3]triazol-2-yl)hexan-1-amine, PANI: polyaniline, CNT: carbon nanotube, poly(SNS-NH 2 ): poly(4-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzenamine), BlmTh: 2-heptyl-4,7-di(thiophen-2-yl)-1H-benzo[d]imidazole, Fmoc-Gly-OH: 2-(((9H-fluoren-9-yl)methoxy)carbonylamino)acetic acid, SNS-COOH: 2-(2,5-di(thiophen-2-yl)-1H-pyrrol-1yl)acetic acid, Lys: lysine, PAMAM: poly(amidoamine) dendrimer (G: generation), GOx: glucose oxidase, ChO: choline oxidase, HRP: horseradish peroxidase, AChE: acetylcholinesterase, ChOx: cholesterol oxidase, and RSD: relative standard deviation.
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Scheme 1. Schematic representation of the proposed HKCN-based biosensors.
The working electrode coated with the conducting polymer layer was used to immobilize either PyOx or G. oxydans. The polymer film was formed on the electrode surface using 10 voltammetric cycles between 1.0 V and − 1.0 V at a scan rate of 0.1 V/s in TBAPF 6 (0.1 M)/dichloromethane medium. For the enzyme immobilization, proper amounts of PyOx solution (0.5 mg in 5.0 μL, 50 mM sodium phosphate buffer, pH 7.0) and glutaraldehyde solution (5.0 μL, 1.0% in sodium phosphate buffer, pH 7.0) were spread over the polymer coated graphite electrode (in the case of FIA mode, a glassy carbon electrode (GCE) was used). Then the electrodes were allowed to stand at ambient conditions for 90 min. In the case of whole cell immobilization, 5.0 μL of G. oxydans cell containing 0.68 × 109 cell titer, suspended in sodium phosphate buffer (50 mM, pH 7.0), was spread over the polymer coated graphite electrode (GCE for FIA) and glutaraldehyde solution (5.0 μL, 1.0% in sodium phosphate buffer, pH 7.0) added and the electrodes were allowed to dry at ambient conditions for 60 min. Daily inoculated cells were used for
biosensor construction. Microbial biosensors containing fresh cells were daily prepared. 2.6. Measurements The basic principle of the measurement was based on the following of oxygen consumption due to the catalytic activity of the enzyme or the respiratory activity of microorganisms (depending on the biological material) in the presence of glucose as a substrate. The decrease in the amount of oxygen was monitored at −0.7 V versus the Ag/AgCl reference electrode. The biosensing system was optimized at batch mode and characterization studies were performed in FIA configuration. In the batch mode, all the measurements were performed under constant magnetic stirring of the solution in the cell containing 10 mL sodium phosphate buffer (50 mM, pH 6.5 or pH 7.0). The three electrodes were immersed into the cell and kept in working buffer solution for 5 min and the working
Fig. 2. Fluorescence images of poly(HKCN) before (A) and after immobilization of PyOx (B) and G. oxydans (B) under optimized conditions on ITO glass (with 40× and 100× magnification for HKCN/PyOx and HKCN/G. oxydans respectively).
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Fig. 3. Effect of pH (A) and scan number of electropolymerization (B) on the biosensor response of HKCN/PyOx biosensor (in sodium phosphate buffer, 50 mM, pH 6.5, −0.7 V). Error bars show S.D. of three measurements for each.
Fig. 4. Calibration curve for glucose in batch mode (A) and FIA mode (B) (in sodium phosphate buffer, 50 mM, pH 7.5, −0.7 V). Error bars show S.D. of three measurements for each. Insets: Time dependent current response with the addition of 1.0 mM glucose for batch and FIA modes.
buffer solution has been refreshed after each measurement. After the current became constant, glucose was added to the reaction cell. Current value was recorded after a steady state current had been achieved. Biosensor response was obtained as the difference between first and second steady state currents. The electrodes were washed with distilled water after each measurement. In the flow injection analysis system, the three electrodes were integrated into a flow through the cell and the measurements were carried out at room temperature in working buffer. Glucose standard solutions and samples were applied as substrates by a computer-controlled injection valve. The obtained current signals were plotted as a calibration curve and glucose concentrations in samples were determined using the calibration curve.
2.7. Sample application The developed biosensors were examined with real samples (coke and fruit juice). Samples were degassed and diluted with working buffer and then injected into the FIA system. Calibration curves were used to determine the glucose contents in measured samples. The samples were also applied to a commercial enzyme assay kit based on spectrophotometric Trinder reaction (Cromatest, Glucose MR, Cat. No. 1129010) as the reference method and results were compared with those obtained with the constructed biosensors. In the Trinder reaction, the glucose is oxidized to D-gluconate by GOx with the formation of hydrogen peroxide. In the presence of peroxidase (POD), a mixture of phenol and 4-
Table 2 Comparison of analytical characteristics of pyranose oxidase based biosensors prepared using various immobilization materials. Electrode
Immobilization method
Linear range
RSD
Operational stability
Reference
CPE
CNT-modified Unmodified Entrapment via AuNPs–PANI/AgCl/gelatin nanocomposite matrix Covalent immobilization Osmium redox polymer type II with poly(ethylene glycol) diglycidyl Osmium redox polymer type I cross linking via diglycidyl ether Osmium redox polymer type II cross linking via diglycidyl ether Electropolymerization of HKCN Electropolymerization of HKCN
0.2–30.0 mM 0.5–30.0 mM 0.05–0.75 mM 40–650 μM 10–400 μM 0.25–6.0 mM 0.125–2.0 mM 0.05–1.0 mM 0.01–1.0 mM
2.3% 2.5% 2.16% 0.3–1.3% 4.4% 2.8% 0.6% 3.5% 4.8%
4% decrease after 7 h (21 measurements) 7% decrease after 7 h (21 measurements) 12% decrease after 6 h 20% decrease after 5 h – 22% decrease after 8 h 6% decrease after 18 h – No decrease after 81 measurements
[41]
GCE CPE Graphite Graphite Graphite Graphite (batch mode) GCE (FIA mode)
[20] [44] [45] [46] [46] This work This work
CPE: carbon paste electrode, CNT: carbon nanotube, GCE: glassy carbon electrode, FIA: flow injection analysis, AuNPs: gold nanoparticles, PANI: polyaniline, and HKCN: 4-amino-N-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzamide.
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Fig. 5. Substrate specificity of biosensor (A) (in sodium phosphate buffer, 50 mM, pH 6.5, −0.7 V). Effects of various interferents to electrode response in the presence of glucose (B) (in sodium phosphate buffer, 50 mM, pH 6.5, −0.7 V).
aminoantipyrine (4-AAP) is oxidized by hydrogen peroxide to form a red quinoneimine dye proportional to the glucose concentration in the sample [32].
3. Results and discussion Various types of biosensors based on conducting polymers have been designed for the analysis of important biological compounds. In previous studies, biological components were either co-deposited with conducting polymers or immobilized to the surfaces immediately after electropolymerization of monomer. Functionalized unique conducting polymers such as 4-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl) benzenamine (SNS-NH2) [24,25,33], poly(6-(4,7-di(thiophen-2-yl)-2Hbenzo[d][1–3]triazol-2-yl)hexan-1-amine) [poly(TBT6–NH2)] [34], and poly(2-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl) (SNS) acetic acid) [28] have been utilized to obtain high stable bio-detection systems through strong bonds between polymeric platform and biomolecules. In this study, a novel conducting polymer, N-(2,5-di(thiophen-2-yl)-1Hpyrrol-1-yl)4-aminobenzamide (HKCN) was used to prepare PyOx and G. oxydans biosensors. The comparison of the analytical performance of conducting polymer based biosensors has been summarized in Table 1. The easy adaptation of conducting polymer based biosensors to FIA mode is the main advantage of this work. In FIA mode, much more analysis can be carried out via on-line monitoring of compounds with computer-based measurements. Synthesis and characterization studies of N-(2,5-di(thiophen-2-yl)1H-pyrrol-1-yl)4-aminobenzamide were performed and published in our previous study [30]. In the present study, PyOx and G. oxydans were used as biological materials. The conducting polymer of HKCN was electrochemically deposited onto a graphite electrode by cyclic voltammetry technique. Electropolymerization of the monomer (in dichloromethane/TBAPF6 (0.1 M) solvent–electrolyte system) by a repeated potential scan at 0.1 V/s on graphite, up to 10 cycles was shown in Fig. 1. After deposition of monomer and the enzyme immobilization step, charge and thickness of the polymer film have been calculated as 425.16 μC and 9.45 nm for the HKCN/PyOx biosensor. In the case of the bacterial biosensor, these values were found as 211.5 μC and 4.7 nm, respectively. Immobilization of biological materials was performed through covalent binding using the glutaraldehyde as a bifunctional agent. Glutaraldehyde forms a covalent bond between amino groups of electropolymerized monomer and amino groups of enzyme or free amino groups on the cell surface (Scheme 1). Poly(HKCN) deposition and immobilization of bio-components were also monitored by fluorescence microscopy after forming on the ITO glass (Fig. 2).
3.1. HKCN/PyOx biosensors PyOx is a flavoprotein with 300,000 kDa molecular weight. There are three key differences between GOx and PyOx. The most important one is its high affinity for D-glucose (Km ∼ 1.0 mM). Another one is that PyOx oxidizes both the α-form as well as the β-form of pyranose as substrates, thus, the addition of mutarotase is not required to detect D-glucose. Additionally, PyOx catalyzes the C-2 oxidation of D-glucose (glucose oxidase catalyzes C-1 oxidation of D-glucose) with high affinity for its corresponding 2-keto sugars with a concomitant generation of H2O2. Its ability to efficiently catalyze the oxidation of several sugars provides the use of PyOx in large-scale industrial applications, biotechnology and biofuel cells in which lignocellulose hydrolysate is oxidized and thus, used for small-scale energy production [20,41]. 3.1.1. Optimization studies In batch mode operations, the effects of several parameters such as pH, cycle number for the electrochemical deposition of the monomer and amount of biological material were optimized. Afterwards, developed biosensors were characterized at FIA mode. First, the effects of pH and cycle number on the sensitivity of the HKCN/PyOx biosensor using glucose were optimized. The pH of the buffer was varied from acidic to neutral (pH 5.0–7.0). Relative biosensor responses (%) were versus tested pHs (Fig. 3A). The maximum response was observed at pH 6.5. The optimum pH for free PyOx was reported as 7.0–7.5 depending on the enzyme sources [42]. This shift on the pH optima could be due to the effect of cationic character of HKCN. To investigate the effect of cycle number, HKCN was deposited onto the surfaces using different scan numbers (10, 20 and 40 cycles). The biosensor in which 20 cycles were used exhibited the highest signal response (Fig. 3B). Increasing of the thickness caused lower charge transfer rates because of the diffusion barrier. On the other hand, when the number of scans was decreased, lower thicknesses are obtained, and the functional groups on the electrode surface may not be enough to attach enough enzyme molecules on it. This may also cause the lower response as well as decreased
Table 3 Results of determination of total glucose using HKCN/PyOx biosensor and spectrophotometric method in beverages. Sample
Glucose (M)a HKCN/PyOx
Spectrophotometric
Recovery %
Coke Fizzy Fizzy (with orange)
0.130 ± 0.016 0.112 ± 0.009 0.175 ± 0.015
0.126 ± 0.006 0.113 ± 0.002 0.163 ± 0.001
103 99 107
a
Results were given as mean ± standard deviation (n: 3).
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(Fig. 5A). The effects of common interfering compounds on the amperometric responses of the prepared HKCN/PyOx biosensor were also studied. The obtained current values were shown in Fig. 5B. Finally, the application of the HKCN/PyOx biosensor to measure the glucose content of real samples such as fruit juice, coke and drinks was carried out. Diluted samples with the working buffer were injected to the reaction medium instead of glucose. The content of real samples was calculated via calibration graph for glucose, and compared with the data obtained with the spectrophotometric method (Table 3). The high recovery values showed that the biosensing analyses were not affected from the nature of the sample matrix. 3.2. HKCN/G. oxydans biosensors
Fig. 6. Effect of pH on the biosensor response (pH 6.0–8.0, sodium phosphate buffer, 50 mM, −0.7 V, room temperature, 1.0 mM glucose). Error bars show S.D. of three measurements.
biomolecule stability on the surface. The results were in agreement with a previous study [34].
3.1.2. Analytical characteristics Fig. 4 shows the calibration curves for glucose (insets show that current change depends on time after glucose addition). The linearity was obtained between 0.05 and 1.0 mM (Fig. 4A) and 0.01–1.0 mM (Fig. 4B) for HKCN/PyOx biosensors in batch and FIA modes, respectively. To test repeatability, the current signals corresponding to 0.5 mM glucose solutions were measured. The standard deviation and variation coefficient (%) were calculated as ±0.019 and 3.5% (n: 10) and ±0.003 and 4.8% (n: 12), in batch and FIA modes respectively. Michaelis–Menten constant has been also calculated as 0.30 and 0.88 mM according to the Lineweaver–Burk diagram for batch and FIA modes [43]. To achieve operational stability 81 measurements were carried out in FIA mode, and no decrease in the activity was observed. The comparison of the analytical performance of PyOx-based biosensors has been given in Table 2. To test the amperometric responses for various monosaccharides, these sugars were injected into the reaction medium instead of glucose. When the biosensor response to glucose was considered as 100%, the responses of HKCN/PyOx biosensor to galactose, xylose, mannose and fructose were calculated as 27%, 31.2%, 28.5%, 32.7%, respectively
3.2.1. Optimization studies The effect of the pH on the biosensor performance was investigated with 1.0 mM glucose, by using 50 mM sodium phosphate buffer systems between pH 6.0 and 8.0. As given in Fig. 6, the best result obtained at pH 7.0, it was chosen as the optimum working pH and used for further experiments. The second parameter investigated was the film thickness, which has a direct effect on the biosensor response because the stability of the polymer layer and response time of the biosensor highly depend on film thickness [47]. The amount of the conducting polymer on the electrode surface can be controlled by adjusting the scan number in electropolymerization process, which enhances homogeneous film formation on the electrode surface, regardless of its shape or size [48]. The polymers were deposited on the graphite working electrode with scans of 5, 10 and 20 cycles respectively. After the modification of each electrode, microbial biosensors were prepared as given in Materials and method. The relationship between the biosensor response and electropolymerization time was shown in Fig. 7A. The best response was recorded, when the polymer was deposited within 10 cycles. The increase in the polymer thickness caused lower current responses. This might be due to the restricted adsorption and diffusion capacities to interact with the target analyte. On the other hand, lower signals were obtained with the decreased film thickness due to lack of enough functional groups of the conducting polymer to keep the cells on the surface. As a result, the electrode prepared with 10 cycles was selected for subsequent experiments. Finally, the effect of cell amount on biosensor response was tested. For this purpose, biosensors containing different amounts of bacterial cells (0.34 × 109, 0.68 × 109 and 1.36 × 109 cell titers, respectively) were prepared and calibrated for the glucose (Fig. 7B). The best signal responses were obtained with the 0.68 × 109 cell titer. Therefore, all measurements were conducted by using this cell amount.
Fig. 7. Effect of deposition time (A) and effect of cell amount on the biosensor response (B) (in sodium phosphate buffer, 50 mM, pH 7.0, −0.7 V). Error bars show S.D. of three measurements.
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Fig. 8. Calibration curve for glucose in batch mode (A) and FIA mode (B) (in sodium phosphate buffer, 50 mM, pH 7.0, −0.7 V). Error bars show S.D. of three measurements for each. Insets: Time dependent current response with the addition of 1.0 mM and 2.5 mM glucose for batch and FIA modes.
3.2.2. Analytical characteristics Analytical characteristics of the microbial biosensor were investigated under the optimized conditions. Linear graphs were plotted for current (μA) versus glucose concentration (mM) for both batch and FIA modes. When the microbial sensor was applied for batch system, the linear range was found to be in the range of 0.25 mM to 2.5 mM, with the equation of y = 0.102x + 0.216 (R2 = 0.997) (Fig. 8A). The response time was 20 s after glucose addition. For the FIA mode, linearity was obtained in the range of 0.1–7.5 mM glucose and described by the equation of y = 0.046x + 0.118 (R2 = 0.998) (Fig. 8B). In this case the response time was 10 s. The FIA combined microbial biosensor showed higher performance compared with the batch system with a corresponding shorter-response time and better linear range. In view of the characteristics of FIA mode, it is an automated system and suitable for the process control owing to its intrinsic characteristics of rapidity, simplicity and versatility [49,50]. Further characterization studies of the proposed microbial system were performed at FIA. During the FIA experiments, the same experimental conditions that were determined in batch conditions such as optimal pH, electropolymerization time and cell amount were used. The repeatability of the system was examined for 2.5 mM glucose by 10 repetitive injections and standard deviations (SD) and coefficient of variation (cv) were calculated as ±0.035 mM and 1.87%, respectively. Furthermore, the reproducibility of 3 electrodes, which were prepared under the same conditions on different days, was also examined. The relative standard deviation (RSD) for three different electrodes was found as 4.6% for 1.0 mM glucose. Additionally, 66 repetitive measurements were done
using 7.5 mM glucose to investigate the operational stability. Only, 11% decrease was observed after this period. According to these findings, it can be said that this system has good repeatability and operational stability in the optimized working conditions. In order to increase the performance of a biosensor, the basic requirement is relevant to usage of a suitable immobilization method [51]. Conducting polymers form suitable environment for biomolecules and the usage of them for biomolecule immobilization provides construction of advanced biosensor devices because of their relative stability, ease of preparation and good conductivity [52,29]. The comparison of the analytical performance of G. oxydansbased biosensors has been summarized in Table 4. Our results reveal that the polymer provides a good platform and enables cells to keep their metabolic activity during the optimum experimental conditions. In order to examine the amperometric responses of HKCN/G. oxydans biosensors to different monosaccharides such as fructose, glucose, xylose and some alcohols such as ethanol and glycerol, they were used as substrates instead of glucose. The obtained results were given in Fig. 9A. To evaluate the selectivity of HKCN/G. oxydans biosensors, the influences of common interference in the detection of glucose were also tested. No current response was observed for ascorbic acid, phenol and uric acid, after that these interfering compounds were injected together with glucose and also no interfering effect of them was obtained for glucose detection (Fig. 9B). The HKCN/G. oxydans biosensor was successfully applied to the determination of glucose in drinks. The obtained results were also compared to results which were found with spectrophotometric method (Table 5). The good agreements were between the obtained values
Table 4 Comparison of analytical characteristics of G. oxydans based biosensors prepared using various immobilization materials. Electrode
Immobilization method
Linear range
RSD
Operational stability
Reference
Graphite Graphite GCE Graphite Graphite GCE (batch mode) GCE (FIA mode) Graphite (batch mode) GCE (FIA mode)
CNT-modified chitosan Electropolymerization of SNS(NO2) CHIT–Fc hybrid Electropolymerization of BDT Electropolymerization of SNS-NH2 Adsorption on TM-Mont Adsorption on TM-Mont Electropolymerization of HKCN Electropolymerization of HKCN
0.05–1.0 mM 0.25–4.0 mM 1.5–25.0 mM 0.5–2.0 mM 0.1–2.5 mM 0.1–5.0 mM 0.15–5.0 mM 0.25–2.5 mM 0.1–7.5 mM
2.3% 4.2% 4.8% 2.85% 2.9% – 3.79% – 1.87%
15% decrease after 5 h 6.0% decrease after 5 h 15% decrease after 3 h 16.7% decrease after 4 h 11% decrease after 5 h – 27.1% decrease after 3 h – 11% decrease after 3 h
[53] [26] [31] [23] [33] [54] [54] This work This work
GCE: glassy carbon electrode, FIA: flow injection analysis, CNT: carbon nanotube, SNS(NO2): poly(1-(4-nitrophenyl)-2,5-di(2-thienyl)-1H-pyrrole), CHIT: chitosan, Fc: ferrocene, BDT: 4,7-di(2,3)-dihydrothienol[3,4-b][1,4]dioxin-5-yl-benzo[1,2,5]thiadiazole, SNS-NH2: 4-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzenamine, TM: trimethylamine, Mont: montmorillonite, and HKCN: 4-amino-N-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzamide.
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Fig. 9. Substrate specificity of biosensor (A) (in sodium phosphate buffer, 50 mM, pH 7.0, −0.7 V). Effects of various interferents to electrode response in the presence of glucose (B) (in sodium phosphate buffer, 50 mM, pH 7.0, −0.7 V).
indicating that HKCN-based biosensors have a great potential in the practical glucose analysis in sample applications. 4. Conclusion Development of effective biosensing systems using suitable immobilization matrix is very crucial. In our previous study, HKCN was synthesized using hydrazide instead of amine to increase product yield and improve properties of the corresponding polymer. Poly(HKCN) is more stable and it has lower band gap and better long-term stability compared with other 2,5-di(2-thienyl)pyrrole (SNS) derivatives [30]. In this article we aimed to utilize a newly synthesized HKCN monomer in biosensor fabrication as an immobilization material and to test them in real samples without any interference. When HKCN-based biosensors are compared with other conducting polymer based biosensors, the main advantage is the easy integration with the FIA system. The proposed systems do not require any complicated immobilization procedure for the construction. The preparation of the biosensor is very simple, cheap and not time consuming. The biosensors showed a good linear range and repeatability. The application of an automatic FIA system is more superior than batch systems for process control because of no feedback to control, reduced waste, no necessity of an expert operator, high speed, and precision. PyOx and G. oxydans are chosen as the model biological components. PyOx efficiently catalyzes the oxidation of several sugars and the ability of PyOx to oxidize various sugars provides the use of PyOx in largescale industrial applications, biotechnology and biofuel cells in which lignocellulose hydrolysate is oxidized and thus used for small-scale energy production [55,56]. It could also be possible to use the proposed biosensors as multianalyte detectors for sugar analysis. Similarly, G. oxydans is particularly suitable for biosensor construction due to some attractive features such as very fast growing in simple cultivation media, high activities, and stability during immobilization, unnecessary with an external cofactor addition to medium for dehydrogenases. Because of the reactive centers of dehydrogenase oriented to the
Table 5 Total glucose analysis using HKCN/G. oxydans biosensor and spectrophotometric method in some beverages. Sample
Fruit juice Coke Fizzy (with orange) a
Glucose (M)a HKCN/G. oxydans
Spectrophotometric
Recovery %
0.131 ± 0.009 0.106 ± 0.0003 0.102 ± 0.003
0.135 ± 0.001 0.111 ± 0.013 0.105 ± 0.007
97 95.5 97
Results were given as mean ± standard deviation (n:3).
periplasmic space, transport of substrates into the cell is not necessary and it simplifies the use of G. oxydans-modified electrodes in biofuel cell applications [57–59]. Briefly, both biosensors were showing good performance in detection of glucose in real samples after characterization studies. Using HKCN based biosensors in FIA mode, 66 and 81 repetitive measurements were carried out with HKCN/G. oxydans and HKCN/PyOx biosensors, respectively. Acknowledgments The authors thank the TUBITAK for the financial support (project nos. 111T074 and 2209—research project supports for university students). References [1] L.D. Mello, L.T. Kubota, Food Chem. 77 (2002) 237–256. [2] R. Monosik, M. Stredansky, J. Tkac, E. Sturdik, Food Anal. Methods 5 (2012) 40–53. [3] V. Velusamy, K. Arshak, O. Korostynska, K. Oliwa, C. Adley, Biotechnol. Adv. 28 (2010) 232–254. [4] M. Nayak, A. Kotian, S. Marathe, D. Chakravortty, Biosens. Bioelectron. 25 (2009) 661–667. [5] P. D'Orazio, Clin. Chim. Acta 412 (2011) 1749–1761. [6] B.D. Malhotra, A. Chaubey, Sensors Actuators B Chem. 91 (2003) 117–127. [7] A. Qureshi, Y. Gurbuz, J.H. Niazia, Sensors Actuators B Chem. (2012) 171–172 (62–76). [8] J.R. Windmiller, J. Wang, Electroanalysis 25 (2013) 29–46. [9] B. Pejcic, R. De Marco, G. Parkinson, Analyst 131 (2006) 1079–1090. [10] T.H. Scheper, J.M. Hilmer, F. Lammers, C. Müller, M. Reinecke, J. Chromatogr. A 725 (1996) 3–12. [11] T. Scheper, J. Ind. Microbiol. 9 (1992) 163–172. [12] M. Akin, A. Prediger, M. Yuksel, T. Höpfner, D. Odaci Demirkol, S. Beutel, S. Timur, T. Scheper, Biosens. Bioelectron. 26 (2011) 4532–4537. [13] S. Rodriguez-Mozaz, M.P. Marco, M.J. Lopez de Alda, D. Barceló, Anal. Bioanal. Chem. 378 (2004) 588–598. [14] A.J. Baeumner, Anal. Bioanal. Chem. 377 (2003) 434–445. [15] I.E. Tothill, Comput. Electron. Agric. 30 (2001) 205–218. [16] S. Kröger, S. Piletsky, A.P.F. Turner, Mar. Pollut. Bull. 45 (2002) 24–34. [17] P.J. Vikesland, K.R. Wigginton, Environ. Sci. Technol. 44 (2010) 3656–3669. [18] M. Yuksel, M. Akın, C. Geyik, D. Odaci Demirkol, C. Ozdemir, A. Bluma, T. Höpfner, S. Beutel, S. Timur, T. Scheper, Biotechnol. Progr. 27 (2011) 530–538. [19] K. Damar, D. Odaci Demirkol, Talanta 87 (2011) 67–73. [20] C. Ozdemir, F. Yeni, D. Odaci, S. Timur, Food Chem. 119 (2010) 380–385. [21] D. Odaci Demirkol, S. Timur, Microchim. Acta 173 (2011) 537–542. [22] N. Yildirim, D. Odaci, G. Ozturk, S. Alp, Y. Ergun, K.H. Feller, K. Dornbusch, S. Timur, Dyes Pigments 89 (2011) 144–148. [23] E. Baskurt, F. Ekiz, D. Odaci Demirkol, S. Timur, L. Toppare, Colloids Surf. B: Biointerfaces 97 (2012) 13–18. [24] C. Ozdemir, S. Tuncagil, D. Odaci Demirkol, S. Timur, L. Toppare, J. Macromol. Sci. Part A Pure Appl. Chem. 48 (2011) 503–508. [25] S. Tuncagil, C. Ozdemir, D. Odaci Demirkol, S. Timur, L. Toppare, Food Chem. 127 (2011) 1317–1322. [26] S. Tuncagil, D. Odaci, S. Varıs, S. Timur, L. Toppare, Bioelectrochemistry 76 (2009) 169–174. [27] K. Arshak, V. Velusamy, O. Korostynska, K. Oliwa-Stasiak, C. Adley, IEEE Sens. J. 9 (2009) 195–1942. [28] S. Demirci, F.B. Emre, F. Ekiz, F. Oguzkaya, S. Timur, C. Tanyeli, L. Toppare, Analyst 137 (2012) 4254–4261. [29] B.D. Malhotra, A. Chaubey, S.P. Singh, Anal. Chim. Acta 578 (2006) 59–74.
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Contents lists available at ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
A novel functional conducting polymer as an immobilization platform Emine Guler a, Hakan Can Soyleyici b, Dilek Odaci Demirkol a,⁎, Metin Ak c, Suna Timur a,⁎ a b c
Ege University, Faculty of Science, Biochemistry Department, 35100 Bornova, Izmir, Turkey Adnan Menderes University, Faculty of Art and Science, Chemistry Department, Aydin, Turkey Pamukkale University, Faculty of Art and Science, Chemistry Department, Denizli, Turkey
a r t i c l e
i n f o
Article history: Received 15 December 2013 Received in revised form 18 February 2014 Accepted 22 March 2014 Available online 31 March 2014 Keywords: Conducting polymers Pyranose oxidase G. oxydans Biosensing
a b s t r a c t Here, we present the fabrication of conducting polymer based enzymatic and microbial biosensors. To obtain immobilization platforms for both pyranose oxidase (PyOx) and Gluconobacter oxydans, the graphite electrode surface was modified with the polymer of 4-amino-N-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzamide (HKCN) which has free amino groups on the surface for further bioconjugation reactions with the biomolecules. Initially, the electrode surface was covered with HKCN via electropolymerization. Then, either PyOx or G. oxydans cell was stabilized using glutaraldehyde as a cross-linker. After optimization of biosensors, analytical characterization and surface imaging studies were investigated. The change of current depends on glucose concentration between 0.05–1.0 mM and 0.25–2.5 mM with HKCN/PyOx and HKCN/G. oxydans biosensors in batch systems. Also, the calibration graphs were obtained for glucose in FIA mode, and in this case, linear ranges were found to be 0.01–1.0 mM and 0.1–7.5 mM for HKCN/PyOx and HKCN/G. oxydans, respectively. © 2014 Elsevier B.V. All rights reserved.
1. Introduction Biomolecule-based detection technology has become a dynamic area of research because of its remarkable potential for a variety of applications such as in food/beverage industries [1–4], diagnosis of disease [5–9], bioprocess monitoring [10–12] and screening of environmental pollutants [13–17]. To improve the stability of the biosensors for different requirements, modification of transducer surface for immobilization of bio-molecule is a key point. Different strategies such as self-assembly monolayers [18,19], natural polymers (gelatin, chitosan etc.) [20,21], sol–gel method [22], and conducting polymers [23–25] have been used to prepare immobilization platforms. The use of conducting polymers in the design of biosensors is very common because of homogeneous and manageable film character, ability of modification of physical and optical properties, stability and biocompatibility, availability of various types of monomers, reproducibility and easy production, and efficient electron transfer ability [26–28]. The synthesis of novel conducting polymers gives the creativity of the fabricated bio-detection systems. And also the addition of functional groups to the monomer backbone provides special and targeted immobilization of bio-components and also improves the stability of bio-components onto conducting polymers because of linkage via covalent attachment [29]. Here we described the use of a novel monomer with amino groups as an immobilization matrix. In the first part, 2,5-di(2-thienyl)pyrrole ⁎ Corresponding authors. Tel.: +90 2323115487; fax: +90 2323115485. E-mail addresses: [email protected], [email protected] (D.O. Demirkol), [email protected], [email protected] (S. Timur).
http://dx.doi.org/10.1016/j.msec.2014.03.063 0928-4931/© 2014 Elsevier B.V. All rights reserved.
derivative namely 4-amino-N-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl) benzamide (HKCN) has been synthesized via the reaction of 1,4-di(2thienyl)-1,4-butanedione and p-aminobenzoyl hydrazide. Using hydrazide instead of amine has not only increased product yield, but also improved properties of the corresponding polymer. Recently, optical properties of HKCN were described [30]. Herein, electrochemically deposited HKCN onto the graphite surfaces was used for the immobilization of PyOx and whole cell of Gluconobacter oxydans to fabricate electrochemical enzyme and whole cell biosensors. After optimization, the characterization studies such as linearity were carried out using both batch and FIA modes.
2. Materials and method 2.1. Reagents Pyranose oxidase (PyOx; pyranose:oxygen 2-oxidoreductase, E.C. 1.1.3.10, from Coriolus sp.), D-glucose, D(+)-xylose, ethanol, glycerol, and ascorbic acid were purchased from Sigma Chem. Co. and tetrabutylammoniumhexafluorophosphate (TBAPF6) was purchased from Aldrich. Dichloromethane (DCM) was obtained from Merck. Glutaraldehyde solution (25%, v/v) was purchased from Sigma-Aldrich. Commercial enzyme assay kit (Glucose MR, Cat. No. 1129010) was obtained from Cromatest. Aluminum chloride (AlCl3) (Aldrich), succinyl chloride (Aldrich), hydrochloric acid (Merck), NaHCO3 (Aldrich), MgSO4 (Aldrich), propionic acid (Aldrich), and toluene (Aldrich) were used for the synthesis of the monomer. All other chemicals were of analytical grade.
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2.2. Apparatus All amperometric measurements were carried out by PalmSens Electrochemical Measurement Unit (Palm Instruments, Houten, Netherlands). During batch operations, experiments were performed in a reaction cell (10 mL) at room temperature using a three electrode configuration, consisting of a graphite working electrode, a Ag/AgCl reference electrode (3.0 M KCl, Metrohm, Switzerland) and a platinum counter electrode (Metrohm, Switzerland). For the flow injection measurements the electrodes were mounted into an electrochemical flow-through cell of the cross-flow type with glassy carbon working, Ag/AgCl reference, and platinum wire counter electrodes (CHI130, Austin, www.chinstruments.com). The FIA system contains a peristaltic pump (FIAtron, Oconomowoc, WI, USA), an eightport injection valve (FIAtron, Oconomowoc, WI) and a cross-flow cell with three electrodes. The peristaltic pump equipped with Teflon tubing (0.50 mm inner diameter) carried 1.6 mL buffer solution per minute. Samples were injected with the eight-port injection valve with a 100 μL sample injection loop. The FIA system was connected to a PalmSens potentiostat for the electrochemical measurements. Colorimetric experiments were performed with a Pharmacia LKB Novaspec II spectrophotometer (LKB Biochrom, England). An Olympus BX53F fluorescence microscope, an Olympus DP72 camera and Uplanapo 100× objective were used in fluorescence images of monomer and prepared biosensors. 2.3. Cell cultivation The strain of G. oxydans DSMZ 2343 was obtained from the German Collection of Microorganisms and Cell Cultures (Braunschweig, Germany) and was sub-cultured on the agar containing (g/L): D-glucose, 100; yeast extract, 10; calcium carbonate, 20; agar, 20. Then, the cells were inoculated into liquid medium of the following composition, g/L: D-glucose, 5.0 and yeast extract, 5.0. Cells were grown at 28 °C on a rotary shaker (175 rpm) in 250 mL flasks with 50 mL medium. Cell growth was followed spectrophotometrically by measuring the optical density at 600 nm. Cells were collected in their late exponential phase, separated from culture medium by centrifugation (4000 rpm, 10 min), washed with sterile 0.9% NaCl solution and re-centrifuged. The obtained cellular paste was used for biosensor preparation [31]. 2.4. Synthesis of HKCN The monomer and its reagents (1,4-di(2-thienyl)-1,4-butanedione and p-aminobenzoyl hydrazide) were prepared according to the
Fig. 1. Repeated potential-scan electropolymerization of monomer in dichloromethane/ TBAPF6 (0.1 M) solvent–electrolyte system at a scan rate of 0.1 V/s on graphite (up to 10 cycles).
methods described in the literature [30]. All experiments were carried out under dry argon by using Standard Schlenk techniques. Solvents were dried, distilled and saturated with argon. A round-bottomed flask equipped with an argon inlet and magnetic stirrer was charged with 2.5 g (10 mmol) 1,4-di(2-thienyl)-1,4butanedione, 1.51 g (10 mmol) p-aminobenzoyl hydrazide, 0.2 g (1.2 mmol) PTSA, 1.0 mL DMSO and 20 mL toluene. The resultant mixture was stirred and refluxed for 18 h under argon. The darkened solution was filtered hot to remove oily decomposition products. Then, the mixture was cooled to room temperature and the solid product was filtered off to give a yellow powder that was washed with pentane (3 × 15 mL) and air-dried (yield: 3.1 g, 85%).
2.5. Preparation of HKCN/PyOx and HKCN/G. oxydans biosensors Before each experiment, graphite rods (Ringsdorff Werke GmbH, Bonn, Germany, 3.05 mm diameter and 13% porosity) were polished with wet emery paper and rinsed thoroughly with distilled water to be used in batch operations. Glassy carbon electrodes (GCEs) were cleaned by polishing with 0.05 μm alumina slurry followed by ultrasonication in ethanol and distilled water for 5 min prior to being used in flow injection mode of analysis (FIA).
Table 1 Comparison of analytical performance of conducting polymer based biosensors. Electrode
Conducting polymer
Biological material
Target analyte
Linearity
RSD
Reference
Graphite Graphite Graphite Graphite GCE Graphite Graphite Graphite Graphite Graphite Graphite
PBDT PESeE BEDOA-6 Poly(TBT6-NH2) PANI–CNT Poly(SNS-NH2) BlmTh:Fmoc-Gly-OH SNS-COOH SNS-COOH/Lys SNS-COOH/PAMAM G2 SNS-COOH/PAMAM G4
GOx GOx GOx ChO HRP AChE and ChO ChOx GOx GOx GOx GOx
Glucose Glucose Glucose Choline H2O2 Acetylcholine Cholesterol Glucose Glucose Glucose Glucose
0.05–2.0 mM 0.01–2.0 mM 0.025–1.25 mM 0.1–10.0 mM 0.2–19.0 μM 0.12–10.0 mM 0.3–10 μM 0.01–1.2 mM 0.01–2.40 mM 0.02–1.20 mM 0.02–1.20 mM
– – 5.87% 3.7% 4.6% 3.2% 4.4% – – – –
[35] [35] [36] [34] [37] [38] [39] [40] [28] [28] [28]
GCE: glassy carbon electrode, PBDT: poly(4,7-di(2,3)-dihydrothienol[3,4-b][1,4]dioxin-5-yl-benzo[1,2,5]thiadiazole), PESeE: poly(4,7-di(2,3)-dihydrothienol[3,4-b][1,4]dioxin-5-yl-2,1,3benzoselenadiazole), BEDOA-6: 6-(4,7-bis(2,3-dihydrothieno[3,4-b][1,4]dioxin-5-yl)-2H-benzo[d][1–3]triazol-2-yl)hexan-1-amine, poly(TBT6-NH2); poly(6-(4,7-di(thiophen-2-yl)2Hbenzo[d][1–3]triazol-2-yl)hexan-1-amine, PANI: polyaniline, CNT: carbon nanotube, poly(SNS-NH 2 ): poly(4-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzenamine), BlmTh: 2-heptyl-4,7-di(thiophen-2-yl)-1H-benzo[d]imidazole, Fmoc-Gly-OH: 2-(((9H-fluoren-9-yl)methoxy)carbonylamino)acetic acid, SNS-COOH: 2-(2,5-di(thiophen-2-yl)-1H-pyrrol-1yl)acetic acid, Lys: lysine, PAMAM: poly(amidoamine) dendrimer (G: generation), GOx: glucose oxidase, ChO: choline oxidase, HRP: horseradish peroxidase, AChE: acetylcholinesterase, ChOx: cholesterol oxidase, and RSD: relative standard deviation.
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Scheme 1. Schematic representation of the proposed HKCN-based biosensors.
The working electrode coated with the conducting polymer layer was used to immobilize either PyOx or G. oxydans. The polymer film was formed on the electrode surface using 10 voltammetric cycles between 1.0 V and − 1.0 V at a scan rate of 0.1 V/s in TBAPF 6 (0.1 M)/dichloromethane medium. For the enzyme immobilization, proper amounts of PyOx solution (0.5 mg in 5.0 μL, 50 mM sodium phosphate buffer, pH 7.0) and glutaraldehyde solution (5.0 μL, 1.0% in sodium phosphate buffer, pH 7.0) were spread over the polymer coated graphite electrode (in the case of FIA mode, a glassy carbon electrode (GCE) was used). Then the electrodes were allowed to stand at ambient conditions for 90 min. In the case of whole cell immobilization, 5.0 μL of G. oxydans cell containing 0.68 × 109 cell titer, suspended in sodium phosphate buffer (50 mM, pH 7.0), was spread over the polymer coated graphite electrode (GCE for FIA) and glutaraldehyde solution (5.0 μL, 1.0% in sodium phosphate buffer, pH 7.0) added and the electrodes were allowed to dry at ambient conditions for 60 min. Daily inoculated cells were used for
biosensor construction. Microbial biosensors containing fresh cells were daily prepared. 2.6. Measurements The basic principle of the measurement was based on the following of oxygen consumption due to the catalytic activity of the enzyme or the respiratory activity of microorganisms (depending on the biological material) in the presence of glucose as a substrate. The decrease in the amount of oxygen was monitored at −0.7 V versus the Ag/AgCl reference electrode. The biosensing system was optimized at batch mode and characterization studies were performed in FIA configuration. In the batch mode, all the measurements were performed under constant magnetic stirring of the solution in the cell containing 10 mL sodium phosphate buffer (50 mM, pH 6.5 or pH 7.0). The three electrodes were immersed into the cell and kept in working buffer solution for 5 min and the working
Fig. 2. Fluorescence images of poly(HKCN) before (A) and after immobilization of PyOx (B) and G. oxydans (B) under optimized conditions on ITO glass (with 40× and 100× magnification for HKCN/PyOx and HKCN/G. oxydans respectively).
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Fig. 3. Effect of pH (A) and scan number of electropolymerization (B) on the biosensor response of HKCN/PyOx biosensor (in sodium phosphate buffer, 50 mM, pH 6.5, −0.7 V). Error bars show S.D. of three measurements for each.
Fig. 4. Calibration curve for glucose in batch mode (A) and FIA mode (B) (in sodium phosphate buffer, 50 mM, pH 7.5, −0.7 V). Error bars show S.D. of three measurements for each. Insets: Time dependent current response with the addition of 1.0 mM glucose for batch and FIA modes.
buffer solution has been refreshed after each measurement. After the current became constant, glucose was added to the reaction cell. Current value was recorded after a steady state current had been achieved. Biosensor response was obtained as the difference between first and second steady state currents. The electrodes were washed with distilled water after each measurement. In the flow injection analysis system, the three electrodes were integrated into a flow through the cell and the measurements were carried out at room temperature in working buffer. Glucose standard solutions and samples were applied as substrates by a computer-controlled injection valve. The obtained current signals were plotted as a calibration curve and glucose concentrations in samples were determined using the calibration curve.
2.7. Sample application The developed biosensors were examined with real samples (coke and fruit juice). Samples were degassed and diluted with working buffer and then injected into the FIA system. Calibration curves were used to determine the glucose contents in measured samples. The samples were also applied to a commercial enzyme assay kit based on spectrophotometric Trinder reaction (Cromatest, Glucose MR, Cat. No. 1129010) as the reference method and results were compared with those obtained with the constructed biosensors. In the Trinder reaction, the glucose is oxidized to D-gluconate by GOx with the formation of hydrogen peroxide. In the presence of peroxidase (POD), a mixture of phenol and 4-
Table 2 Comparison of analytical characteristics of pyranose oxidase based biosensors prepared using various immobilization materials. Electrode
Immobilization method
Linear range
RSD
Operational stability
Reference
CPE
CNT-modified Unmodified Entrapment via AuNPs–PANI/AgCl/gelatin nanocomposite matrix Covalent immobilization Osmium redox polymer type II with poly(ethylene glycol) diglycidyl Osmium redox polymer type I cross linking via diglycidyl ether Osmium redox polymer type II cross linking via diglycidyl ether Electropolymerization of HKCN Electropolymerization of HKCN
0.2–30.0 mM 0.5–30.0 mM 0.05–0.75 mM 40–650 μM 10–400 μM 0.25–6.0 mM 0.125–2.0 mM 0.05–1.0 mM 0.01–1.0 mM
2.3% 2.5% 2.16% 0.3–1.3% 4.4% 2.8% 0.6% 3.5% 4.8%
4% decrease after 7 h (21 measurements) 7% decrease after 7 h (21 measurements) 12% decrease after 6 h 20% decrease after 5 h – 22% decrease after 8 h 6% decrease after 18 h – No decrease after 81 measurements
[41]
GCE CPE Graphite Graphite Graphite Graphite (batch mode) GCE (FIA mode)
[20] [44] [45] [46] [46] This work This work
CPE: carbon paste electrode, CNT: carbon nanotube, GCE: glassy carbon electrode, FIA: flow injection analysis, AuNPs: gold nanoparticles, PANI: polyaniline, and HKCN: 4-amino-N-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzamide.
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Fig. 5. Substrate specificity of biosensor (A) (in sodium phosphate buffer, 50 mM, pH 6.5, −0.7 V). Effects of various interferents to electrode response in the presence of glucose (B) (in sodium phosphate buffer, 50 mM, pH 6.5, −0.7 V).
aminoantipyrine (4-AAP) is oxidized by hydrogen peroxide to form a red quinoneimine dye proportional to the glucose concentration in the sample [32].
3. Results and discussion Various types of biosensors based on conducting polymers have been designed for the analysis of important biological compounds. In previous studies, biological components were either co-deposited with conducting polymers or immobilized to the surfaces immediately after electropolymerization of monomer. Functionalized unique conducting polymers such as 4-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl) benzenamine (SNS-NH2) [24,25,33], poly(6-(4,7-di(thiophen-2-yl)-2Hbenzo[d][1–3]triazol-2-yl)hexan-1-amine) [poly(TBT6–NH2)] [34], and poly(2-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl) (SNS) acetic acid) [28] have been utilized to obtain high stable bio-detection systems through strong bonds between polymeric platform and biomolecules. In this study, a novel conducting polymer, N-(2,5-di(thiophen-2-yl)-1Hpyrrol-1-yl)4-aminobenzamide (HKCN) was used to prepare PyOx and G. oxydans biosensors. The comparison of the analytical performance of conducting polymer based biosensors has been summarized in Table 1. The easy adaptation of conducting polymer based biosensors to FIA mode is the main advantage of this work. In FIA mode, much more analysis can be carried out via on-line monitoring of compounds with computer-based measurements. Synthesis and characterization studies of N-(2,5-di(thiophen-2-yl)1H-pyrrol-1-yl)4-aminobenzamide were performed and published in our previous study [30]. In the present study, PyOx and G. oxydans were used as biological materials. The conducting polymer of HKCN was electrochemically deposited onto a graphite electrode by cyclic voltammetry technique. Electropolymerization of the monomer (in dichloromethane/TBAPF6 (0.1 M) solvent–electrolyte system) by a repeated potential scan at 0.1 V/s on graphite, up to 10 cycles was shown in Fig. 1. After deposition of monomer and the enzyme immobilization step, charge and thickness of the polymer film have been calculated as 425.16 μC and 9.45 nm for the HKCN/PyOx biosensor. In the case of the bacterial biosensor, these values were found as 211.5 μC and 4.7 nm, respectively. Immobilization of biological materials was performed through covalent binding using the glutaraldehyde as a bifunctional agent. Glutaraldehyde forms a covalent bond between amino groups of electropolymerized monomer and amino groups of enzyme or free amino groups on the cell surface (Scheme 1). Poly(HKCN) deposition and immobilization of bio-components were also monitored by fluorescence microscopy after forming on the ITO glass (Fig. 2).
3.1. HKCN/PyOx biosensors PyOx is a flavoprotein with 300,000 kDa molecular weight. There are three key differences between GOx and PyOx. The most important one is its high affinity for D-glucose (Km ∼ 1.0 mM). Another one is that PyOx oxidizes both the α-form as well as the β-form of pyranose as substrates, thus, the addition of mutarotase is not required to detect D-glucose. Additionally, PyOx catalyzes the C-2 oxidation of D-glucose (glucose oxidase catalyzes C-1 oxidation of D-glucose) with high affinity for its corresponding 2-keto sugars with a concomitant generation of H2O2. Its ability to efficiently catalyze the oxidation of several sugars provides the use of PyOx in large-scale industrial applications, biotechnology and biofuel cells in which lignocellulose hydrolysate is oxidized and thus, used for small-scale energy production [20,41]. 3.1.1. Optimization studies In batch mode operations, the effects of several parameters such as pH, cycle number for the electrochemical deposition of the monomer and amount of biological material were optimized. Afterwards, developed biosensors were characterized at FIA mode. First, the effects of pH and cycle number on the sensitivity of the HKCN/PyOx biosensor using glucose were optimized. The pH of the buffer was varied from acidic to neutral (pH 5.0–7.0). Relative biosensor responses (%) were versus tested pHs (Fig. 3A). The maximum response was observed at pH 6.5. The optimum pH for free PyOx was reported as 7.0–7.5 depending on the enzyme sources [42]. This shift on the pH optima could be due to the effect of cationic character of HKCN. To investigate the effect of cycle number, HKCN was deposited onto the surfaces using different scan numbers (10, 20 and 40 cycles). The biosensor in which 20 cycles were used exhibited the highest signal response (Fig. 3B). Increasing of the thickness caused lower charge transfer rates because of the diffusion barrier. On the other hand, when the number of scans was decreased, lower thicknesses are obtained, and the functional groups on the electrode surface may not be enough to attach enough enzyme molecules on it. This may also cause the lower response as well as decreased
Table 3 Results of determination of total glucose using HKCN/PyOx biosensor and spectrophotometric method in beverages. Sample
Glucose (M)a HKCN/PyOx
Spectrophotometric
Recovery %
Coke Fizzy Fizzy (with orange)
0.130 ± 0.016 0.112 ± 0.009 0.175 ± 0.015
0.126 ± 0.006 0.113 ± 0.002 0.163 ± 0.001
103 99 107
a
Results were given as mean ± standard deviation (n: 3).
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(Fig. 5A). The effects of common interfering compounds on the amperometric responses of the prepared HKCN/PyOx biosensor were also studied. The obtained current values were shown in Fig. 5B. Finally, the application of the HKCN/PyOx biosensor to measure the glucose content of real samples such as fruit juice, coke and drinks was carried out. Diluted samples with the working buffer were injected to the reaction medium instead of glucose. The content of real samples was calculated via calibration graph for glucose, and compared with the data obtained with the spectrophotometric method (Table 3). The high recovery values showed that the biosensing analyses were not affected from the nature of the sample matrix. 3.2. HKCN/G. oxydans biosensors
Fig. 6. Effect of pH on the biosensor response (pH 6.0–8.0, sodium phosphate buffer, 50 mM, −0.7 V, room temperature, 1.0 mM glucose). Error bars show S.D. of three measurements.
biomolecule stability on the surface. The results were in agreement with a previous study [34].
3.1.2. Analytical characteristics Fig. 4 shows the calibration curves for glucose (insets show that current change depends on time after glucose addition). The linearity was obtained between 0.05 and 1.0 mM (Fig. 4A) and 0.01–1.0 mM (Fig. 4B) for HKCN/PyOx biosensors in batch and FIA modes, respectively. To test repeatability, the current signals corresponding to 0.5 mM glucose solutions were measured. The standard deviation and variation coefficient (%) were calculated as ±0.019 and 3.5% (n: 10) and ±0.003 and 4.8% (n: 12), in batch and FIA modes respectively. Michaelis–Menten constant has been also calculated as 0.30 and 0.88 mM according to the Lineweaver–Burk diagram for batch and FIA modes [43]. To achieve operational stability 81 measurements were carried out in FIA mode, and no decrease in the activity was observed. The comparison of the analytical performance of PyOx-based biosensors has been given in Table 2. To test the amperometric responses for various monosaccharides, these sugars were injected into the reaction medium instead of glucose. When the biosensor response to glucose was considered as 100%, the responses of HKCN/PyOx biosensor to galactose, xylose, mannose and fructose were calculated as 27%, 31.2%, 28.5%, 32.7%, respectively
3.2.1. Optimization studies The effect of the pH on the biosensor performance was investigated with 1.0 mM glucose, by using 50 mM sodium phosphate buffer systems between pH 6.0 and 8.0. As given in Fig. 6, the best result obtained at pH 7.0, it was chosen as the optimum working pH and used for further experiments. The second parameter investigated was the film thickness, which has a direct effect on the biosensor response because the stability of the polymer layer and response time of the biosensor highly depend on film thickness [47]. The amount of the conducting polymer on the electrode surface can be controlled by adjusting the scan number in electropolymerization process, which enhances homogeneous film formation on the electrode surface, regardless of its shape or size [48]. The polymers were deposited on the graphite working electrode with scans of 5, 10 and 20 cycles respectively. After the modification of each electrode, microbial biosensors were prepared as given in Materials and method. The relationship between the biosensor response and electropolymerization time was shown in Fig. 7A. The best response was recorded, when the polymer was deposited within 10 cycles. The increase in the polymer thickness caused lower current responses. This might be due to the restricted adsorption and diffusion capacities to interact with the target analyte. On the other hand, lower signals were obtained with the decreased film thickness due to lack of enough functional groups of the conducting polymer to keep the cells on the surface. As a result, the electrode prepared with 10 cycles was selected for subsequent experiments. Finally, the effect of cell amount on biosensor response was tested. For this purpose, biosensors containing different amounts of bacterial cells (0.34 × 109, 0.68 × 109 and 1.36 × 109 cell titers, respectively) were prepared and calibrated for the glucose (Fig. 7B). The best signal responses were obtained with the 0.68 × 109 cell titer. Therefore, all measurements were conducted by using this cell amount.
Fig. 7. Effect of deposition time (A) and effect of cell amount on the biosensor response (B) (in sodium phosphate buffer, 50 mM, pH 7.0, −0.7 V). Error bars show S.D. of three measurements.
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Fig. 8. Calibration curve for glucose in batch mode (A) and FIA mode (B) (in sodium phosphate buffer, 50 mM, pH 7.0, −0.7 V). Error bars show S.D. of three measurements for each. Insets: Time dependent current response with the addition of 1.0 mM and 2.5 mM glucose for batch and FIA modes.
3.2.2. Analytical characteristics Analytical characteristics of the microbial biosensor were investigated under the optimized conditions. Linear graphs were plotted for current (μA) versus glucose concentration (mM) for both batch and FIA modes. When the microbial sensor was applied for batch system, the linear range was found to be in the range of 0.25 mM to 2.5 mM, with the equation of y = 0.102x + 0.216 (R2 = 0.997) (Fig. 8A). The response time was 20 s after glucose addition. For the FIA mode, linearity was obtained in the range of 0.1–7.5 mM glucose and described by the equation of y = 0.046x + 0.118 (R2 = 0.998) (Fig. 8B). In this case the response time was 10 s. The FIA combined microbial biosensor showed higher performance compared with the batch system with a corresponding shorter-response time and better linear range. In view of the characteristics of FIA mode, it is an automated system and suitable for the process control owing to its intrinsic characteristics of rapidity, simplicity and versatility [49,50]. Further characterization studies of the proposed microbial system were performed at FIA. During the FIA experiments, the same experimental conditions that were determined in batch conditions such as optimal pH, electropolymerization time and cell amount were used. The repeatability of the system was examined for 2.5 mM glucose by 10 repetitive injections and standard deviations (SD) and coefficient of variation (cv) were calculated as ±0.035 mM and 1.87%, respectively. Furthermore, the reproducibility of 3 electrodes, which were prepared under the same conditions on different days, was also examined. The relative standard deviation (RSD) for three different electrodes was found as 4.6% for 1.0 mM glucose. Additionally, 66 repetitive measurements were done
using 7.5 mM glucose to investigate the operational stability. Only, 11% decrease was observed after this period. According to these findings, it can be said that this system has good repeatability and operational stability in the optimized working conditions. In order to increase the performance of a biosensor, the basic requirement is relevant to usage of a suitable immobilization method [51]. Conducting polymers form suitable environment for biomolecules and the usage of them for biomolecule immobilization provides construction of advanced biosensor devices because of their relative stability, ease of preparation and good conductivity [52,29]. The comparison of the analytical performance of G. oxydansbased biosensors has been summarized in Table 4. Our results reveal that the polymer provides a good platform and enables cells to keep their metabolic activity during the optimum experimental conditions. In order to examine the amperometric responses of HKCN/G. oxydans biosensors to different monosaccharides such as fructose, glucose, xylose and some alcohols such as ethanol and glycerol, they were used as substrates instead of glucose. The obtained results were given in Fig. 9A. To evaluate the selectivity of HKCN/G. oxydans biosensors, the influences of common interference in the detection of glucose were also tested. No current response was observed for ascorbic acid, phenol and uric acid, after that these interfering compounds were injected together with glucose and also no interfering effect of them was obtained for glucose detection (Fig. 9B). The HKCN/G. oxydans biosensor was successfully applied to the determination of glucose in drinks. The obtained results were also compared to results which were found with spectrophotometric method (Table 5). The good agreements were between the obtained values
Table 4 Comparison of analytical characteristics of G. oxydans based biosensors prepared using various immobilization materials. Electrode
Immobilization method
Linear range
RSD
Operational stability
Reference
Graphite Graphite GCE Graphite Graphite GCE (batch mode) GCE (FIA mode) Graphite (batch mode) GCE (FIA mode)
CNT-modified chitosan Electropolymerization of SNS(NO2) CHIT–Fc hybrid Electropolymerization of BDT Electropolymerization of SNS-NH2 Adsorption on TM-Mont Adsorption on TM-Mont Electropolymerization of HKCN Electropolymerization of HKCN
0.05–1.0 mM 0.25–4.0 mM 1.5–25.0 mM 0.5–2.0 mM 0.1–2.5 mM 0.1–5.0 mM 0.15–5.0 mM 0.25–2.5 mM 0.1–7.5 mM
2.3% 4.2% 4.8% 2.85% 2.9% – 3.79% – 1.87%
15% decrease after 5 h 6.0% decrease after 5 h 15% decrease after 3 h 16.7% decrease after 4 h 11% decrease after 5 h – 27.1% decrease after 3 h – 11% decrease after 3 h
[53] [26] [31] [23] [33] [54] [54] This work This work
GCE: glassy carbon electrode, FIA: flow injection analysis, CNT: carbon nanotube, SNS(NO2): poly(1-(4-nitrophenyl)-2,5-di(2-thienyl)-1H-pyrrole), CHIT: chitosan, Fc: ferrocene, BDT: 4,7-di(2,3)-dihydrothienol[3,4-b][1,4]dioxin-5-yl-benzo[1,2,5]thiadiazole, SNS-NH2: 4-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzenamine, TM: trimethylamine, Mont: montmorillonite, and HKCN: 4-amino-N-(2,5-di(thiophen-2-yl)-1H-pyrrol-1-yl)benzamide.
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Fig. 9. Substrate specificity of biosensor (A) (in sodium phosphate buffer, 50 mM, pH 7.0, −0.7 V). Effects of various interferents to electrode response in the presence of glucose (B) (in sodium phosphate buffer, 50 mM, pH 7.0, −0.7 V).
indicating that HKCN-based biosensors have a great potential in the practical glucose analysis in sample applications. 4. Conclusion Development of effective biosensing systems using suitable immobilization matrix is very crucial. In our previous study, HKCN was synthesized using hydrazide instead of amine to increase product yield and improve properties of the corresponding polymer. Poly(HKCN) is more stable and it has lower band gap and better long-term stability compared with other 2,5-di(2-thienyl)pyrrole (SNS) derivatives [30]. In this article we aimed to utilize a newly synthesized HKCN monomer in biosensor fabrication as an immobilization material and to test them in real samples without any interference. When HKCN-based biosensors are compared with other conducting polymer based biosensors, the main advantage is the easy integration with the FIA system. The proposed systems do not require any complicated immobilization procedure for the construction. The preparation of the biosensor is very simple, cheap and not time consuming. The biosensors showed a good linear range and repeatability. The application of an automatic FIA system is more superior than batch systems for process control because of no feedback to control, reduced waste, no necessity of an expert operator, high speed, and precision. PyOx and G. oxydans are chosen as the model biological components. PyOx efficiently catalyzes the oxidation of several sugars and the ability of PyOx to oxidize various sugars provides the use of PyOx in largescale industrial applications, biotechnology and biofuel cells in which lignocellulose hydrolysate is oxidized and thus used for small-scale energy production [55,56]. It could also be possible to use the proposed biosensors as multianalyte detectors for sugar analysis. Similarly, G. oxydans is particularly suitable for biosensor construction due to some attractive features such as very fast growing in simple cultivation media, high activities, and stability during immobilization, unnecessary with an external cofactor addition to medium for dehydrogenases. Because of the reactive centers of dehydrogenase oriented to the
Table 5 Total glucose analysis using HKCN/G. oxydans biosensor and spectrophotometric method in some beverages. Sample
Fruit juice Coke Fizzy (with orange) a
Glucose (M)a HKCN/G. oxydans
Spectrophotometric
Recovery %
0.131 ± 0.009 0.106 ± 0.0003 0.102 ± 0.003
0.135 ± 0.001 0.111 ± 0.013 0.105 ± 0.007
97 95.5 97
Results were given as mean ± standard deviation (n:3).
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