Biosensors and Bioelectronics 57 (2014) 199–206
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Label-free electrochemical immunosensor based on gold–silicon carbide nanocomposites for sensitive detection of human chorionic gonadotrophin Long Yang a,1, Hui Zhao b,1, Shuangmei Fan a, Shuangsheng Deng b, Qi Lv b, Jie Lin b,n, Can-Peng Li a,n a b
School of Chemical Science and Technology, Yunnan University, Kunming 650091, PR China School of Life Science, Yunnan University, Kunming 650091, PR China
art ic l e i nf o
a b s t r a c t
Article history: Received 12 November 2013 Received in revised form 8 February 2014 Accepted 10 February 2014 Available online 19 February 2014
Uniform and highly dispersed gold–silicon carbide (Au@SiC) nanocomposites were prepared via simple way and used for fabrication of label-free electrochemical immunosensor for sensitive detection of human chorionic gonadotrophin (hCG). Using Au@SiC as electrode material and using ferricyanide as mediator, the proposed immunosensor provides a simple and economic method with higher sensitivity and a wider concentration range for detection of hCG. Under the optimal condition, the approach provided a good linear response range from 0.1 to 5 IU/L and 5 to 1000 IU/L with a low detection limit of 0.042 IU/L. The immunosensor showed good selectivity, acceptable stability and reproducibility. Satisfactory results were obtained for determination of hCG in human serum samples. The proposed method provides a promising platform of clinical immunoassay for other biomolecules. In addition, the bio-functionalization of SiC combined with other nanomaterials will provide promising approach for electrochemical sensing and biosensing platform. & 2014 Elsevier B.V. All rights reserved.
Keywords: Gold–silicon carbide nanocomposite Human chorionic gonadotrophin Label-free Electrochemical immunosensor
1. Introduction Human chorionic gonadotrophin (hCG) is a secretion of the placenta during pregnancy and gestational trophoblastic diseases (Marcillac et al., 1992). It is increased as a consequence of abnormal placental invasion and placental immaturity. Therefore, it is an important diagnostic marker of pregnancy and one of the most important carbohydrate tumor markers (Lu et al., 2012). Thus, exactly determining the concentration of hCG in urine or serum plays an important role in monitoring of trophoblastic diseases in all modern immunological pregnancy tests (Yang et al., 2011a). Specific affinity between antibody and corresponding antigen, so-called immunoassay, provides a promising analytical method for clinical assay and biochemical analysis (Mani et al., 2009; Tang and Ren, 2008; Cui et al., 2007; Nourani et al., 2013). During recent years, conventional diagnostic methods, such as enzyme-linked immunosorbent assay, chemiluminescence, surface plasmon resonance, and quartz crystal microbalance, have been the main
n
Corresponding authors. Tel./fax: þ86 871 6503 1119. E-mail addresses: [email protected] (J. Lin), [email protected] (C.-P. Li). 1 These two authors contributed equally to this work.
http://dx.doi.org/10.1016/j.bios.2014.02.019 0956-5663 & 2014 Elsevier B.V. All rights reserved.
methods used for detection hCG (Zhou et al., 2010; Krishnan et al., 2011; Miller et al., 2011; Yang et al., 2010). Compared with conventional immunoassays, electrochemical immunoassay has exhibited several advantages, including simplicity of instrument, low cost, feasibility of miniaturization, and subsequent portability. Some strategies based on sandwich-type immunosensors (Yang et al., 2009; Wei et al., 2011; Viet et al., 2013) have been applied to the determination of hCG. In comparison to sandwich-type immunosensors, label-free immunosensors have obvious advantages that its fabrication is simple, easy-handle, and low-cost owing to its avoidance of tedious labeling operations (Wu et al., 2013). Up to now, various label-free immunosensors related to medical diagnosis (Zhuo et al., 2008; Song et al., 2010), environmental monitoring (Tran et al., 2012), and food safety monitoring (Li et al., 2011a) have been reported. Among them, electrochemical label-free immunosensors have been attracted increasing attention due to its high sensitivity, low cost, and ease of preparation. In the design and fabrication of highly sensitive electrochemical immunosensors, antibody immobilization and signal amplification are the crucial steps (Sánchez et al., 2008). Many kinds of nanomaterial, including noble metal nanoparticles, carbon nanomaterials, semiconductor nanoparticles, metal oxide nanostructures, and hybrid nanostructures, have been developed to amplify
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electrochemical signal in order to improve the sensitivity of electrochemical immunosensor (Pei et al., 2013). Gold nanoparticles (Au NPs) are well-known bio-nanomaterials because of their large specific surface area, strong adsorption ability, well suitability, and good conductivity (Liu et al., 2005); it can strongly interact with biomaterials and has been utilized as an intermediator to immobilize antibody to efficiently retain its activity and to enhance current response in the construction of immunosensor (Huang et al., 2013). Chitosan (CS) is a polysaccharide derived by deacetylation of chitin. It possesses many advantages, such as excellent membrane-forming ability, high permeability towards water, good adhesion, and biocompatibility. Also, it has abundant reactive amino and hydroxyl functional groups. Therefore, it has been widely used as an immobilization matrix for biofabrication. Ferricyanide is an excellent redox mediator in the electrochemical immunosensor system. It can cause a lower background current and have a pair of redox peaks in the amperometric measurement. The electrochemical signal achieved by ferricyanide is very stable. Thus, using ferricyanide as mediator is beneficial for improving the stability of the immunosensor. Silicon carbide (SiC) is a material that consists of the covalent bonding of Si and C atoms, in a tetrahedron form in which Si (or C) is the central atom. The high mechanical and chemical stability of the material are determined by the very short bond length, and hence, a very high bond strength present in the SiC structure (Deva Reddy et al., 2008). SiC has more than 200 polymorphic forms, called polytypes, but cubic (β)3C-SiC, hexagonal 4H-SiC, and (α)6H-SiC are the most common polytypes (Oliveros et al., 2013). As a kind of electronic matrices and a wide band gap semiconductor, SiC has been demonstrated attractive properties, such as high modulus, high strength, good corrosion/oxidation resistance, and good high-temperature strength (Belmonte et al., 2006; Van Dorp et al., 2009; Ferroni and Pezzotti 2002; Willander et al., 2006). Hence, SiC has great practical application in several scopes, such as catalysis oxidize, photocatalyst reaction, selective oxidation of hydrogen sulfide into elemental sulfur, and isomerization of linear saturated hydrocarbons (Salimi et al., 2009; Wu et al., 2011; Dai et al., 2012). However, the design and fabrication of electrochemical immunosensor based on SiC and its nanocomposite has not been developed. Most biomolecule recognition-based systems require immobilization of specific molecules with controlled structural order and composition (Oliveros et al., 2013). Surface functionalization provides many advantages in the development of semiconductor based biosensors. In addition, surface functionalization is one of the main tools used for covalent biomolecule immobilization. SiC is a very promising and interesting material for surface functionalization because the formation of a very thin native oxide on its surface facilitates the successful surface termination that is the prerequisite in the realization of devices. However, the surface functionalization of SiC with carboxyl group has not been explored. It has been reported that if SiC is appropriately doped, the conductivity of this material dramatically increases and exhibits electrical characteristics similar to carbon materials (Chu et al., 1995). In the work performed by Wu et al., they were able to resolve the overlapping voltammetric responses of ascorbic acid (AA), dopamine (DA) and uric acid (UA) on a SiC-coated glassy carbon electrode (GCE), and the selective determination of DA in the presence of AA and UA with a high sensitivity (Wu et al., 2011). Salimi et al. used SiC nanoparticles to modify a GCE to detect insulin via electrocatalytic oxidation with high sensitivity, excellent catalytic activity, short response time, and long term stability (Salimi et al., 2009). These findings led to the construction of different SiC electrode applications. In the present paper, gold–silicon carbide (Au@SiC) nanocomposites were prepared via simple way and used for fabrication of
label-free electrochemical immunosensor. Using Au@SiC as electrode material and using ferricyanide as mediator, the proposed label-free immunosensor provides a simple and economic method with higher sensitivity and a wider concentration range for detection of hCG and could find potential application in clinical analysis.
2. Materials and methods 2.1. Reagents and apparatus SiC was purchased from Nanjing Aipurei Nano-Material Company (Nanjing, China). Montmorillonite and kaolin (KL) were purchased from Shanghai Chemical Reagent Co. Ltd. (Shanghai, China). SiO2 was prepared from montmorillonite according to a previous report (Yang et al., 2013). Anti-hCG, hCG, luteinizing hormone (LH), thyroid stimulating hormone (TSH), and follicle-stimulating hormone (FSH) were obtained from National Institutes for Food and Drug Control (Beijing, China). Bovine serum albumin (BSA), gold chloride (HAuCl4), and chitosan (CS), were obtained from Sigma Chemical Co. (St. Louis, MO, USA). All other reagents were of analytical grade. Phosphate buffer (PBS, 0.1 M, pH 7.4) containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl was used as working solution. All aqueous solutions were prepared with deionized water (18 MΩ/cm). Cyclic voltammetry (CV) experiments were performed with a CHI 660E Electrochemical Workstation from Shanghai Chenhua Instrument (Shanghai, China) and conducted using a threeelectrode system, with the proposed immunosensor as working electrode, a platinum wire as the counter electrode, a saturated calomel electrode (SCE) as the reference electrode. SiC and Au@SiC were characterized by a QUNT200 scanning electron microscopy (SEM, USA), a JEM 2100 transmission electron microscopy (TEM, Japan), a Rigaku TTR III X-ray diffractometer (XRD, Japan), and a Thermo Fisher SCIENTIFIC Nicolet IS10 Fourier transform infrared spectrometry (FTIR, USA). The size of Au@SiC NPs was determined by dynamic light scattering (DLS) using a Zetasizer Nano-ZS90 (Malvern Instruments, UK). 2.2. Preparation of Au@SiC SiC was carboxyl functionalized as previously reported (An et al., 2007) with the modification of replacing SiO2 with SiC before using.
Scheme 1. Schematic illustration of the stepwise immunosensor fabrication process.
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Au@SiC nanocomposite was prepared as follows. Briefly, SiC–COOH (5 mg), polyethylene glycol 400 (0.1 ml), and HAuCl4 solution (0.01 M, 0.5 ml) were dispersed into 20 ml of deionized water, and then the mixture was stirred with a magnetic stirrer for 24 h at room temperature. One milliliter of 50 mM ascorbic acid solution
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was added dropwise and stirred for 1 h and shaken for an additional 12 h. After centrifuging and washing with deionized water for three times, the resulting Au@SiC nanocomposite was redispersed in 0.2% CS solution to obtain Au@SiC–CS suspension (resulting concentration 0.5 mg/ml) for subsequent use.
Fig. 1. SEM images of SiC (A) and Au@SiC (B); TEM images of SiC (C) and Au@SiC (D); The size distribution histogram of Au@SiC (E); Nitrogen adsorption–desorption isotherms of the SiC (F); FTIR spectra of SiC and SiC–COOH (G); XRD patterns of SiC and Au@SiC (H).
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2.3. Fabrication of the electrochemical immunosensor
2.4. Electrochemical measurements
Glassy carbon electrode (GCE, 3 mm in diameter) was polished with 0.3 and 0.05 mm Al2O3 powder respectively and subsequently sonicated in ethanol and deionized water to remove the physically adsorbed substance and dried in air. To prepare the Au@SiC modified electrode, 5 ml of the Au@SiC dispersed by 0.2% CS was dropped onto the electrode surface and dried at room temperature. Then, the Au@SiC–CS modified electrode was submerged in anti-hCG solution at 4 1C for 12 h to yield anti-hCG/Au@SiC–CS/GCE. Finally, to block possible remaining active sites and eliminate the risk of nonspecific binding, 0.25% BSA dissolved by PBS (0.1 M, pH 7.4) was coated on the electrode for 30 min. After every modificatory step, the modified electrode was cleaned with PBS (0.1 M, pH 7.4) to remove the physically absorbed species. To visualize the self-assembled process of the immunosensor, a schematic illustration was given in Scheme 1.
The immunoreaction was performed by immersing the immunosensor in 0.1 M pH 7.4 PBS containing various concentrations of hCG at 33 1C for 30 min. Subsequently, the proposed electrode was rinsed with PBS to remove the residue and unbound antigen from the electrode. After the specific reaction of antibody–antigen, the formed antigen–antibody immunocomplex on the electrode surface hindered the electron transfer toward the electrode surface, resulting in a decrease of electrochemical signal. Thus, the quantitative detection of hCG could be accomplished by tracking the electrical signals. Differential pulse voltammetry (DPV) was applied in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl from 0.05 to 0.5 V with a pulse amplitude of 50 mV and a pulse width of 50 ms.
3. Results and discussion 3.1. Characterization of Au@SiC
Fig. 2. CVs of bare GCE (a), CS/GCE (b), KL–CS/GCE (c), SiO2–CS/GCE (d), SiC–CS/GCE (e), Au@SiC–CS/GCE (f), anti-hCG/Au@SiC–CS/GCE (g), BSA/anti-hCG/Au@SiC–CS/ GCE (h), and hCG/BSA/anti-hCG/Au@SiC–CS/GCE (i) in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl. Scan rate: 0.05 V/s at 25 1C (A); DPVs of bare GCE (a), CS/GCE (b), KL–CS/GCE (c), SiO2–CS/GCE (d), SiC–CS/GCE (e), Au@SiC–CS/ GCE (f), anti-hCG/Au@SiC–CS/GCE (g), BSA/anti-hCG/Au@SiC–CS/GCE (h), and hCG/ BSA/anti-hCG/Au@SiC–CS/GCE (i) in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 / 4 and 0.1 M KCl. Pulse width: 0.05s; amplitude: 0.05V (B).
The morphologies and microstructures of SiC and Au@SiC were investigated by SEM and TEM observation. As shown in Fig. 1(A for SEM and C for TEM), SiC NPs demonstrated large amounts of particles about 45 nm. Compared with Fig. 1(A and C), Fig. 1(B for SEM and D for TEM) showed that large amounts of Au NPs as spherical particles about 10 nm were reduced on SiC NPs, indicating the successful preparation of Au@SiC nanocomposites. The size of Au@SiC NPs was determined by DLS using a Malvern Zetasizer Nano-ZS90. From the size distribution histogram of as shown in Fig. 1E, an average size of 370 nm for Au@SiC was obtained. Since the size of SiC and Au NPs demonstrated by TEM was 45 nm and 10 nm, respectively. It can be concluded that the 370 nm size for Au@SiC determined by DLS was the aggregation of several Au@SiC NPs. Fig. 1F showed the low-temperature nitrogen adsorption– desorption isotherms of SiC, which gave a clear IV type adsorptive isothermal curves, the adsorptive capacity rise sharply when the relative pressure is in the range from 0.80 to 0.95, and the hysteresis loop existed at the desorption process. As we known, the shapes of the isotherms and the hysteresis loop suggested the sample possesses a mesoporous structure. Additionally, above the relative pressure of 0.9, the isotherm still rise, indicating the sample also possesses some macro porous structures. The BET surface areas of the SiC was 40.0 m2/g, which was two times bigger than previously reported (Dai et al., 2012). The results illustrated that the SiC could be suitable for anchoring metal nanoparticles. FTIR spectra was employed to investigate carboxyl functionalization of SiC. Fig. 1G showed FTIR spectra of SiC and SiC–COOH. As observed, SiC displayed a strong band at 829 cm 1, which was ascribed to the typical stretching vibration of Si–C. The two weak bands at 3460 and 1640 cm 1 could be attributed to the stretching vibration and bending vibration of O–H from absorbed water. Compared with that of SiC, the FTIR spectrum of SiC–COOH displayed the presence of C ¼O (1724 cm 1), which confirmed the presence of a carboxyl group. The crystal structure of the Au@SiC nanocomposite was investigated by XRD. Fig. 1H showed the XRD patterns of SiC and Au@SiC nanocomposite. XRD patterns well defined peaks at 35.64, 41.34, 60.03, 71.82, and 75.33 1 (2θ), indicating the formation of the cubic phase of SiC NPs. According to the cubic phase crystal structure of SiC, we can conclude that the crystal form of the SiC is β-SiC (3C–SiC). XRD patterns well defined peaks at 44.35, 64.53, 77.57, and 81.86 o (2θ), indicating the formation of the cubic phase of Au NPs. The characteristic diffraction peak of Au NPs confirmed the presence of Au NPs in the Au@SiC nanocomposites.
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3.2. Electrochemical characterization of immunosensor To characterize the fabrication process of the immunosensor, CVs at all immobilization steps were recorded. As shown in Fig. S1, the Au@SiC–CS film modified GCE showed cathodic current at about –0.65 V in 0.1 M pH7.4 PBS. There is no peak current in the present sweep range from –0.3 V to 0.7 V. This may be attributed to the synergistic effect between Au NPs and SiC. As shown in Fig. 2A, a pair of typical reversible redox peaks of ferricyanide ions can be observed on the bare GCE (curve a) in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl. When a film of CS was immobilized onto a bare GCE, the current response increased obviously due to the fact that CS can facilitate electron transfer (Sassolas et al., 2012) between the solution and the electrode (curve b). When KL–CS, SiO2–CS, and SiC–CS were immobilized onto a bare GCE, SiC–CS/GCE (curve e) showed higher current response than KL–CS/GCE (curve c) and SiO2–CS/GCE (curve d), indicating that SiC is a good promoter of electron transport, and favored for electron communication between the solution and the electrode, which is consistent with previously described (Dai et al., 2012; Wright and Horsfall, 2007; Oliveros et al., 2013). Furthermore, the excellent adsorptive capacity of SiC can also lead to the higher current response. After Au NPs were reduced on SiC, the peak current response (curve f) increased drastically, improving the electron transfer from the redox probe to the electrode. The significantly enhanced current response could be attributed to the fact that Au NPs with splendid conductivity and large surface area could amplify the electrochemical signal (Wu et al., 2013). When anti-hCG was adsorbed onto the electrode through cross-link with the amino group of CS and carboxyl group of SiC, there was an obvious decrease of the current response (curves g), which suggested that antibody had been successfully immobilized on the electrode surface. Moreover, Au NPs with good biocompatibility
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were used to immobilize anti-hCG, Au NPs in the Au@SiC nanocomposites played a crucial role in antibody immobilization. The reason for the decrease of the current response is that anti-hCG as protein biomacromolecules insulated the conductive support and played a key role as obstructions in the process of electron transfer (Fan et al., 2013). While BSA was added onto the electrode, the current response was decreased further for the same reason (curves h). Additionally, the peak current (curve i) further decreased after incubation with 500 IU/L hCG, indicating the formed immunocomplex after specific immunoreaction acted as an electron-transfer and mass-transfer blocking layer, which greatly inhibited the electron transfer toward the electrode surface (Sun et al., 2013). From these results, it could be seen that the Au@SiC–CS film modified electrode could be used for the detection of hCG. The electrochemical behaviors of the fabrication process of the immunosensor were confirmed by DPV. The similar electrochemical characterizations were obtained as shown in Fig. 2B. The kinetics of the electrode reactions was investigated by studying the effect of scan rate at the BSA/anti-hCG/Au@SiC–CS immunoelectrode in 0.1 M pH 7.4 PBS containing 2 mM [Fe (CN)6]3 /4 and 0.1 M KCl. As shown in Fig. 3A, both anodic peak current (Ipa) and cathodic peak current (Ipc) increased with the increase of scan rate in the range of 10–100 mV/s. Also, the Ipa and Ipc showed a linear relationship with the square root of the scan rate respectively (Fig. 3B), suggesting that the electrode reaction is a diffusion-controlled electrochemical process, which is the ideal case for quantitative measurements (Mobin et al., 2010). It can be seen that both Ipa and Ipc of the immunoelectrode increased linearly and were proportional to the square root of the scan rate according to Eqs. Ipa (mA)¼ 9.05v1/2 (mV/s)1/2 þ 10.2 and Ipc (mA)¼10.4v1/2 (mV/s)1/2–13.9. The separation of peaks suggests that the process is not perfectly reversible; however, stable redox peak current and position during repeated scans at a particular
Fig. 3. CV studies of BSA/anti-hCG/Au@SiC–CS/GCE immunoelectrode as a function of scan rate (10–100 mV/s) (A); Magnitude of current response vs. square root of scan rate (B); Value of redox potential vs. scan rate (C); Value of redox potential vs. the Napierian logarithm of scan rate (D).
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scan rate suggests that the BSA/anti-hCG/Au@SiC–CS immunoelectrode exhibits a quasi-reversible process (Topoglidis et al., 2003). Moreover, both anodic peak potential (Epa) and cathodic peak potential (Epc) demonstrated a linear relationship with the scan rate respectively (Fig. 3C), indicating that the electron transport form redox moieties to the electrode is very facile (Kaushik et al., 2013). Epa and Epc were proportional to the scan rate according to equations. Epa (V)¼2.42v (mV/s) þ 0.206 and Epc (V) ¼ 2.87v (mV/s) þ 0.162. Fig. 3D showed the relationship between the peak potentials and the Napierian logarithm of scan rate (ln v). It can be observed that the scan rate affects the position of the redox peaks. With increasing the scan rate, the Epa shifted positively and the Epc shifted negatively. When v is small, both the anodic and cathodic peaks become closed to reversible peaks (Yin et al., 2010). However, the reversibility of electrode reaction changes with the scan rate. The peak separation (ΔEp) may enlarge as the scan rate becomes higher, which indicates the irreversibility of the electrode process increases. The fact that the ΔEp increases with increasing the scanning rate can be attributed to the polarization of the electrodes (Li et al., 2011b). Similarly, a linear relationship between peak potentials and Napierian logarithm of scan rate (ln v) is also observed in the scan rates ranging from 40 to 100 mV/s. The apparent heterogeneous electron transfer rate constant (ks) for the BSA/anti-hCG/Au@SiC–CS immunoelectrode was estimated using the Laviron model (Laviron, 1979) as ks ¼mnFv/RT, where m is peak-to-peak separation, F is Faraday constant, v is scan rate (mV/s), n is the number of transferred electrons and R is gas constant. The value of ks obtained as 3.9 s 1 (T ¼298 K, n ¼1, m ¼0.096 V and v ¼100 mV) is higher than that of other nanoparticles based bioelectrodes (Zhang et al., 2005; Zhao et al., 2005), indicating fast electron transfer between immobilized biomolecules and electrode. 3.3. Optimization of experimental conditions To achieve the optimal immunoassay performance, the pH of the detection solution, the immunoreaction temperature, and immunoreaction time were optimized as important factors that influenced the sensitivity of the proposed immunosensor. The optimization of experimental conditions was supplied in Supplementary Materials. 3.4. Calibration curve of immunosensor
Fig. 4. DPVs of the immunosensor incubated with different concentrations of hCG standard solution (from a to j): 0.0, 1.0, 10.0, 50.0, 100.0, 200.0, 300.0, 500.0, 700.0, and 1000.0 IU/L in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl. Pulse width: 0.05 s; amplitude: 0.05 V (A); Calibration plots of the reduction current change (ΔI) versus concentration of hCG under optimal conditions (B); Selectivity of the proposed immunosensor with FSH (20 IU/L), LH (20 IU/L), TSH (20 mIU/L), hCG (5 IU/L), and hCG (5 IU/L) containing the above mixture of three interferents with the same concentrations (C).
Under the optimal detection condition, the immunosensor was subjected to standard hCG solution with various concentrations. As indicated in Fig. 4A, the peak currents decreased with the increase of hCG concentration, which may be ascribed to that more antigen– antibody immunocomplex as an insulating layer was formed on the surface of electrode, and thus inhibiting the electron transfer. Therefore, the current change was proportional to the concentration of the corresponding hCG. The corresponding calibration curve was plotted in Fig. 4B. The current change was linearly related to hCG concentration in the range of 0.1–5 IU/L and 5–1000 IU/L. Correspondingly, the regression equations were ΔI (mA)¼0.81C (IU/L)þ 2.05 and ΔI (mA)¼ 0.036C (IU/L)þ6.11 with correlation coefficients of 0.996 and 0.998. From the slope of 0.81 the detection limit was calculated to be 0.042 IU/L (signal/noise [S/N]¼3). Compared with other sensors reported previously (Chai et al., 2008; Wang et al., 2010; Yang et al., 2011a; 2011b), the proposed immunosensor exhibited a satisfactory detection limit and linear range. The characteristics of other hCG sensors are summarized in Table 1. Lu et al. reported a sandwich-type immunosensor based on Au-multiwalled carbon nanotubes-graphene composite modified GCE with a very low detection limit of 0.0026 IU/L. However, the fabrication process
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Table 1 Comparison of different immunosensors for detection of hCG. Electrode
Method
Liner range (IU/L)
Detection limit (IU/L)
Ref
anti-hCG/nano-Au/MBa/GCE anti-hCG/nano-gold and CS hybrid film/GCE anti-hCG/Pt–Au alloy nanotube array/GCE HRPb-anti-hCG-hCG modified electrode HRP-anti-hCG/hCG/SGc/GCE Gold-labeled-anti-hCG-Ag(I)-hydroquinone HRP-anti-hCG /hCG/AuNPs/MPSd/GCE HRP-anti-hCG/sol–gel/GEe anti-hCG/GNPsf/pPAg/MWCNTsh/GCE anti-hCG/gold nanotubes array/GCE anti-hCG/Au/MWCNTs/GSi/GCE anti-hCG/AuNPs-TiO2/Thij/GAk/MWCNTs-CS/GCE anti-hCG/Au@SiC-CS/GCE
CV CAl CAl DPV DPV RS assayl DPV DPV CV CAm DPV DPV DPV
1–100 0.2–100 25–400 2.5–12.5 0.5–5.0; 5.0–30 2.5–208.3 5.0–30 0.5–50 1–10, 10–160 0.1–100 0.005–500 0.2–300 0.1–5, 5–1000
0.3 0.1 12 1.4 0.3 0.83 1.4 0.3 0.3 0.08 0.0026 0.08 0.042
Chai et al. (2008) Yang et al. (2009) Tao et al. (2011) Chen et al. (2005) Chen et al. (2006a) Liang et al. (2008) Chen et al. (2006b) Tan et al. (2007) Wang et al. (2010) Yang et al. (2011b) Lu et al. (2012) Yang et al. (2011a) This work
a
Methylene blue. Horseradish peroxidase. c Titania sol–gel matrix. d 3-Mercaptopropanesulfonic acid. e Graphite electrode. f Gold nanoparticles. g Poly-(2,6-pyridinediamine). h Multiwalled carbon nanotubes. i Graphene nanosheets. j Thionine. k Glutaraldehyde. l Chronoamperometry. m Resonance scattering assay. b
increased the complexity of the operation system. In the present work, such a high sensitivity of the proposed immunosensor may be attributed to three factors: (1) SiC’s good biocompatibility, electrical, mechanical, and thermal properties determine its suitability as a biomaterial and biosensing substrate (Wright and Horsfall, 2007); (2) Au NPs possess splendid conductivity, large surface area, and good biocompatibility, which could amplify the electrochemical signal, resulting in high sensitivity, and enhance the stability of the immunosensor; (3) the large amounts of antibodies have been immobilized onto the electrode by Au NPs, the amino group of CS, and carboxyl group of SiC. Also, CS produced a thin film with good biocompatibility and high chemical stability, which increased the sensitivity of the immunosensor.
3.5. Selectivity, reproducibility, and stability It has been reported that some potential interferents such as carcinoembryonic antigen, α-1-fetoprotein, prostate-specific antigen, cancer antigen 125, carbohydrate antigen 19-9, ferritin, BSA, ascorbic acid, L-cysteine, and L-glutamic acid have no interference to the detection of hCG (Wang et al., 2010; Yang et al., 2011a; Tan et al., 2007; Lu et al., 2012; Wu et al., 2013). Thus, three sister molecules of hCG, namely, FSH (20 IU/L), LH (20 IU/L), and TSH (20 mIU/L), which are the known interferents in the hCG detection system, were selected as the predominant interferents in this study. Selectivity was investigated by recording current changes of the proposed immunosensor to the three potential interferents. The result was expressed in Fig. 4C. Compared to the result obtained from the only hCG (5 IU/L), the change of current response of the immunosensor before and after incubation with FSH, LH, and TSH could be neglected. These demonstrated that the three interferents did not cause observable interference to hCG detection, which was attributed to the highly specific antigen–antibody immunoreactions. Additionally, the cross-selectivity of the immunosensor incubated with mixture consisting of hCG and three interferents was also investigated. No remarkable change of ΔI obtained from mixture was observed in comparison with that in the presence of only hCG.
These results indicated that the proposed immunosensor specifically recognized hCG antigen and exhibited good selectivity. The reproducibility of the immunosensor was examined by six equally proposed immunosensors incubated with the same concentration hCG (100 IU/L). The six electrodes exhibited the similar electrochemical responses and a relative standard deviation (RSD) of 3.8% was obtained, indicating satisfying reproducibility. Successive cyclic potential scans for 50 cycles and long-term storage assay were used to examine the stability of the proposed immunosensor. A 4.4% decrease of initial peak current was found after 50 continuous cycle scans. Additionally, the long-term stability experiment was carried out intermittently (every 5 days). When the immunosensor was not in use, it was stored in a refrigerator at 4 1C. Over 95.6% and 88.5% of initial response remained after storage of 15 and 30 days, respectively. The acceptable stability of the immunosensor may be due to the facts that Au NPs loaded on the electrode surface could make protein molecules more firmly attached on electrode. CS as dispersant of Au@SiC nanocomposite had outstanding film-forming ability and high chemical stability. 3.6. Real sample analysis In order to evaluate the feasibility of the proposed immunosenor for real sample analysis, the immunosensor was used for the determination of hCG by standard addition methods in serum samples. The result showed that the recovery was from 98.7% to 103.6% and the RSD was from 1.7% to 7.0% (Table S1), which confirmed the practical value of the immunosensor. In the present work, a simple and effective label-free electrochemical immunosensor based on Au@SiC–CS film was developed. The proposed immunosensor provides a simple and economic method with higher sensitivity and a wider concentration range for detection of hCG. Thus, we believe that there are many opportunities for SiC as an active material to be incorporated in such devices that require certain chemical, mechanical, and electrical properties, which could find potential application in clinical analysis. In addition, the bio-functionalization of SiC
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combined with other nanocomposite and the ongoing research toward biosensor configurations that employ SiC as an active material will increase the possibilities to create complex devices that can perform multiple biomolecule detection and analysis on a single platform. Thus, great care must be taken when profiling multiple analytes simultaneously. 4. Conclusion In summary, the uniform and highly dispersed Au@SiC nanocomposites were synthesized by a facile ascorbic acid assisted chemical reduction. A simple and effective label-free electrochemical immunosensor based on Au@SiC–CS film for hCG detection was developed. The immunosensor displayed a linear response for detection hCG within a wide range (0.1–1000 IU/L). The proposed immunosensor shows a low detection limit (0.042 IU/L), good reproducibility, selectivity, and acceptable stability. The simple fabrication procedure and the ultrasensitivity demonstrated by the immunosensor may provide many potential applications for the detection of hCG in clinical diagnostics. Acknowledgments This work was supported by the Natural Science Foundation of China (31160334; 31260226), the Natural Science Foundation of Yunnan Province (2012FB112; 2012FB121), and the Scientific Research Fund of the Yunnan Provincial Education Department (2011Z051) People’s Republic of China. Appendix A. Supplementary material Supplementary data associated with this article can be found in the online version at http://dx.doi.org/10.1016/j.bios.2014.02.019. References An, Y.Q., Chen, M., Xue, Q.J., Liu, W.M., 2007. J. Colloid Interf. Sci. 311, 507–513. Belmonte, M., Nieto, M.I., Osendi, M.I., Miranzo, P., 2006. J. Eur. Ceram. Soc. 26, 1273–1279. Chai, R., Yuan, R., Chai, Y.Q., Ou, C.F., Cao, S.R., Li, X.L., 2008. Talanta 74, 1330–1336. Chen, J., Yan, F., Dai, Z., Ju, H.X., 2005. Biosens. Bioelectron. 21, 330–336. Chen, J., Tang, J.H., Yan, F., Ju, H.X., 2006a. Biomaterials 27, 2313–2321. Chen, J., Yan, F., Tan, F., Ju, H.X., 2006b. Electroanalysis 18, 1696–1702. Chu, V., Conde, J.P., Jarego, J., Brogueira, P., Rodriguez, J., Barradas, N., et al., 1995. J. Appl. Phys. 78, 3164–3173. Cui, R.J., Pan, H.C., Zhu, J.J., Chen, H.Y., 2007. Anal. Chem. 79, 8494–8501. Dai, H., Chen, Y.L., Lina, Y.Y., Xu, G.F., Yang, C.P., Tong, Y.J., Guo, L.H., Chen, G.N., 2012. Electrochim. Acta 85, 644–649. Deva Reddy, J., Volinsky, A.A., Frewin, C.L., Locke, C., Saddow, S.E., 2008. In Materials Research Society Meeting, ed. Mechanical properties of single and polycrystalline SiC thin films, pp. AA03–AA6. Fan, H.X., Zhang, Y., Wu, D., Ma, H.M., Li, X.J., Li, Y., Wang, H., Li, H., Du, B., Wei, Q., 2013. Anal. Chim. Acta 770, 62–67. Ferroni, L.P., Pezzotti, G., 2002. J. Am. Ceram. Soc. 85, 2033–2038. Huang, K.J., Li, J., Wu, Y.Y., Liu, Y.M., 2013. Bioelectrochemistry 90, 18–23. Kaushik, A., Vasudev, A., Arya, S.K., Bhansali, S., 2013. Biosens. Bioelectron. 50, 35–41.
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Contents lists available at ScienceDirect
Biosensors and Bioelectronics journal homepage: www.elsevier.com/locate/bios
Label-free electrochemical immunosensor based on gold–silicon carbide nanocomposites for sensitive detection of human chorionic gonadotrophin Long Yang a,1, Hui Zhao b,1, Shuangmei Fan a, Shuangsheng Deng b, Qi Lv b, Jie Lin b,n, Can-Peng Li a,n a b
School of Chemical Science and Technology, Yunnan University, Kunming 650091, PR China School of Life Science, Yunnan University, Kunming 650091, PR China
art ic l e i nf o
a b s t r a c t
Article history: Received 12 November 2013 Received in revised form 8 February 2014 Accepted 10 February 2014 Available online 19 February 2014
Uniform and highly dispersed gold–silicon carbide (Au@SiC) nanocomposites were prepared via simple way and used for fabrication of label-free electrochemical immunosensor for sensitive detection of human chorionic gonadotrophin (hCG). Using Au@SiC as electrode material and using ferricyanide as mediator, the proposed immunosensor provides a simple and economic method with higher sensitivity and a wider concentration range for detection of hCG. Under the optimal condition, the approach provided a good linear response range from 0.1 to 5 IU/L and 5 to 1000 IU/L with a low detection limit of 0.042 IU/L. The immunosensor showed good selectivity, acceptable stability and reproducibility. Satisfactory results were obtained for determination of hCG in human serum samples. The proposed method provides a promising platform of clinical immunoassay for other biomolecules. In addition, the bio-functionalization of SiC combined with other nanomaterials will provide promising approach for electrochemical sensing and biosensing platform. & 2014 Elsevier B.V. All rights reserved.
Keywords: Gold–silicon carbide nanocomposite Human chorionic gonadotrophin Label-free Electrochemical immunosensor
1. Introduction Human chorionic gonadotrophin (hCG) is a secretion of the placenta during pregnancy and gestational trophoblastic diseases (Marcillac et al., 1992). It is increased as a consequence of abnormal placental invasion and placental immaturity. Therefore, it is an important diagnostic marker of pregnancy and one of the most important carbohydrate tumor markers (Lu et al., 2012). Thus, exactly determining the concentration of hCG in urine or serum plays an important role in monitoring of trophoblastic diseases in all modern immunological pregnancy tests (Yang et al., 2011a). Specific affinity between antibody and corresponding antigen, so-called immunoassay, provides a promising analytical method for clinical assay and biochemical analysis (Mani et al., 2009; Tang and Ren, 2008; Cui et al., 2007; Nourani et al., 2013). During recent years, conventional diagnostic methods, such as enzyme-linked immunosorbent assay, chemiluminescence, surface plasmon resonance, and quartz crystal microbalance, have been the main
n
Corresponding authors. Tel./fax: þ86 871 6503 1119. E-mail addresses: [email protected] (J. Lin), [email protected] (C.-P. Li). 1 These two authors contributed equally to this work.
http://dx.doi.org/10.1016/j.bios.2014.02.019 0956-5663 & 2014 Elsevier B.V. All rights reserved.
methods used for detection hCG (Zhou et al., 2010; Krishnan et al., 2011; Miller et al., 2011; Yang et al., 2010). Compared with conventional immunoassays, electrochemical immunoassay has exhibited several advantages, including simplicity of instrument, low cost, feasibility of miniaturization, and subsequent portability. Some strategies based on sandwich-type immunosensors (Yang et al., 2009; Wei et al., 2011; Viet et al., 2013) have been applied to the determination of hCG. In comparison to sandwich-type immunosensors, label-free immunosensors have obvious advantages that its fabrication is simple, easy-handle, and low-cost owing to its avoidance of tedious labeling operations (Wu et al., 2013). Up to now, various label-free immunosensors related to medical diagnosis (Zhuo et al., 2008; Song et al., 2010), environmental monitoring (Tran et al., 2012), and food safety monitoring (Li et al., 2011a) have been reported. Among them, electrochemical label-free immunosensors have been attracted increasing attention due to its high sensitivity, low cost, and ease of preparation. In the design and fabrication of highly sensitive electrochemical immunosensors, antibody immobilization and signal amplification are the crucial steps (Sánchez et al., 2008). Many kinds of nanomaterial, including noble metal nanoparticles, carbon nanomaterials, semiconductor nanoparticles, metal oxide nanostructures, and hybrid nanostructures, have been developed to amplify
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electrochemical signal in order to improve the sensitivity of electrochemical immunosensor (Pei et al., 2013). Gold nanoparticles (Au NPs) are well-known bio-nanomaterials because of their large specific surface area, strong adsorption ability, well suitability, and good conductivity (Liu et al., 2005); it can strongly interact with biomaterials and has been utilized as an intermediator to immobilize antibody to efficiently retain its activity and to enhance current response in the construction of immunosensor (Huang et al., 2013). Chitosan (CS) is a polysaccharide derived by deacetylation of chitin. It possesses many advantages, such as excellent membrane-forming ability, high permeability towards water, good adhesion, and biocompatibility. Also, it has abundant reactive amino and hydroxyl functional groups. Therefore, it has been widely used as an immobilization matrix for biofabrication. Ferricyanide is an excellent redox mediator in the electrochemical immunosensor system. It can cause a lower background current and have a pair of redox peaks in the amperometric measurement. The electrochemical signal achieved by ferricyanide is very stable. Thus, using ferricyanide as mediator is beneficial for improving the stability of the immunosensor. Silicon carbide (SiC) is a material that consists of the covalent bonding of Si and C atoms, in a tetrahedron form in which Si (or C) is the central atom. The high mechanical and chemical stability of the material are determined by the very short bond length, and hence, a very high bond strength present in the SiC structure (Deva Reddy et al., 2008). SiC has more than 200 polymorphic forms, called polytypes, but cubic (β)3C-SiC, hexagonal 4H-SiC, and (α)6H-SiC are the most common polytypes (Oliveros et al., 2013). As a kind of electronic matrices and a wide band gap semiconductor, SiC has been demonstrated attractive properties, such as high modulus, high strength, good corrosion/oxidation resistance, and good high-temperature strength (Belmonte et al., 2006; Van Dorp et al., 2009; Ferroni and Pezzotti 2002; Willander et al., 2006). Hence, SiC has great practical application in several scopes, such as catalysis oxidize, photocatalyst reaction, selective oxidation of hydrogen sulfide into elemental sulfur, and isomerization of linear saturated hydrocarbons (Salimi et al., 2009; Wu et al., 2011; Dai et al., 2012). However, the design and fabrication of electrochemical immunosensor based on SiC and its nanocomposite has not been developed. Most biomolecule recognition-based systems require immobilization of specific molecules with controlled structural order and composition (Oliveros et al., 2013). Surface functionalization provides many advantages in the development of semiconductor based biosensors. In addition, surface functionalization is one of the main tools used for covalent biomolecule immobilization. SiC is a very promising and interesting material for surface functionalization because the formation of a very thin native oxide on its surface facilitates the successful surface termination that is the prerequisite in the realization of devices. However, the surface functionalization of SiC with carboxyl group has not been explored. It has been reported that if SiC is appropriately doped, the conductivity of this material dramatically increases and exhibits electrical characteristics similar to carbon materials (Chu et al., 1995). In the work performed by Wu et al., they were able to resolve the overlapping voltammetric responses of ascorbic acid (AA), dopamine (DA) and uric acid (UA) on a SiC-coated glassy carbon electrode (GCE), and the selective determination of DA in the presence of AA and UA with a high sensitivity (Wu et al., 2011). Salimi et al. used SiC nanoparticles to modify a GCE to detect insulin via electrocatalytic oxidation with high sensitivity, excellent catalytic activity, short response time, and long term stability (Salimi et al., 2009). These findings led to the construction of different SiC electrode applications. In the present paper, gold–silicon carbide (Au@SiC) nanocomposites were prepared via simple way and used for fabrication of
label-free electrochemical immunosensor. Using Au@SiC as electrode material and using ferricyanide as mediator, the proposed label-free immunosensor provides a simple and economic method with higher sensitivity and a wider concentration range for detection of hCG and could find potential application in clinical analysis.
2. Materials and methods 2.1. Reagents and apparatus SiC was purchased from Nanjing Aipurei Nano-Material Company (Nanjing, China). Montmorillonite and kaolin (KL) were purchased from Shanghai Chemical Reagent Co. Ltd. (Shanghai, China). SiO2 was prepared from montmorillonite according to a previous report (Yang et al., 2013). Anti-hCG, hCG, luteinizing hormone (LH), thyroid stimulating hormone (TSH), and follicle-stimulating hormone (FSH) were obtained from National Institutes for Food and Drug Control (Beijing, China). Bovine serum albumin (BSA), gold chloride (HAuCl4), and chitosan (CS), were obtained from Sigma Chemical Co. (St. Louis, MO, USA). All other reagents were of analytical grade. Phosphate buffer (PBS, 0.1 M, pH 7.4) containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl was used as working solution. All aqueous solutions were prepared with deionized water (18 MΩ/cm). Cyclic voltammetry (CV) experiments were performed with a CHI 660E Electrochemical Workstation from Shanghai Chenhua Instrument (Shanghai, China) and conducted using a threeelectrode system, with the proposed immunosensor as working electrode, a platinum wire as the counter electrode, a saturated calomel electrode (SCE) as the reference electrode. SiC and Au@SiC were characterized by a QUNT200 scanning electron microscopy (SEM, USA), a JEM 2100 transmission electron microscopy (TEM, Japan), a Rigaku TTR III X-ray diffractometer (XRD, Japan), and a Thermo Fisher SCIENTIFIC Nicolet IS10 Fourier transform infrared spectrometry (FTIR, USA). The size of Au@SiC NPs was determined by dynamic light scattering (DLS) using a Zetasizer Nano-ZS90 (Malvern Instruments, UK). 2.2. Preparation of Au@SiC SiC was carboxyl functionalized as previously reported (An et al., 2007) with the modification of replacing SiO2 with SiC before using.
Scheme 1. Schematic illustration of the stepwise immunosensor fabrication process.
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Au@SiC nanocomposite was prepared as follows. Briefly, SiC–COOH (5 mg), polyethylene glycol 400 (0.1 ml), and HAuCl4 solution (0.01 M, 0.5 ml) were dispersed into 20 ml of deionized water, and then the mixture was stirred with a magnetic stirrer for 24 h at room temperature. One milliliter of 50 mM ascorbic acid solution
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was added dropwise and stirred for 1 h and shaken for an additional 12 h. After centrifuging and washing with deionized water for three times, the resulting Au@SiC nanocomposite was redispersed in 0.2% CS solution to obtain Au@SiC–CS suspension (resulting concentration 0.5 mg/ml) for subsequent use.
Fig. 1. SEM images of SiC (A) and Au@SiC (B); TEM images of SiC (C) and Au@SiC (D); The size distribution histogram of Au@SiC (E); Nitrogen adsorption–desorption isotherms of the SiC (F); FTIR spectra of SiC and SiC–COOH (G); XRD patterns of SiC and Au@SiC (H).
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2.3. Fabrication of the electrochemical immunosensor
2.4. Electrochemical measurements
Glassy carbon electrode (GCE, 3 mm in diameter) was polished with 0.3 and 0.05 mm Al2O3 powder respectively and subsequently sonicated in ethanol and deionized water to remove the physically adsorbed substance and dried in air. To prepare the Au@SiC modified electrode, 5 ml of the Au@SiC dispersed by 0.2% CS was dropped onto the electrode surface and dried at room temperature. Then, the Au@SiC–CS modified electrode was submerged in anti-hCG solution at 4 1C for 12 h to yield anti-hCG/Au@SiC–CS/GCE. Finally, to block possible remaining active sites and eliminate the risk of nonspecific binding, 0.25% BSA dissolved by PBS (0.1 M, pH 7.4) was coated on the electrode for 30 min. After every modificatory step, the modified electrode was cleaned with PBS (0.1 M, pH 7.4) to remove the physically absorbed species. To visualize the self-assembled process of the immunosensor, a schematic illustration was given in Scheme 1.
The immunoreaction was performed by immersing the immunosensor in 0.1 M pH 7.4 PBS containing various concentrations of hCG at 33 1C for 30 min. Subsequently, the proposed electrode was rinsed with PBS to remove the residue and unbound antigen from the electrode. After the specific reaction of antibody–antigen, the formed antigen–antibody immunocomplex on the electrode surface hindered the electron transfer toward the electrode surface, resulting in a decrease of electrochemical signal. Thus, the quantitative detection of hCG could be accomplished by tracking the electrical signals. Differential pulse voltammetry (DPV) was applied in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl from 0.05 to 0.5 V with a pulse amplitude of 50 mV and a pulse width of 50 ms.
3. Results and discussion 3.1. Characterization of Au@SiC
Fig. 2. CVs of bare GCE (a), CS/GCE (b), KL–CS/GCE (c), SiO2–CS/GCE (d), SiC–CS/GCE (e), Au@SiC–CS/GCE (f), anti-hCG/Au@SiC–CS/GCE (g), BSA/anti-hCG/Au@SiC–CS/ GCE (h), and hCG/BSA/anti-hCG/Au@SiC–CS/GCE (i) in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl. Scan rate: 0.05 V/s at 25 1C (A); DPVs of bare GCE (a), CS/GCE (b), KL–CS/GCE (c), SiO2–CS/GCE (d), SiC–CS/GCE (e), Au@SiC–CS/ GCE (f), anti-hCG/Au@SiC–CS/GCE (g), BSA/anti-hCG/Au@SiC–CS/GCE (h), and hCG/ BSA/anti-hCG/Au@SiC–CS/GCE (i) in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 / 4 and 0.1 M KCl. Pulse width: 0.05s; amplitude: 0.05V (B).
The morphologies and microstructures of SiC and Au@SiC were investigated by SEM and TEM observation. As shown in Fig. 1(A for SEM and C for TEM), SiC NPs demonstrated large amounts of particles about 45 nm. Compared with Fig. 1(A and C), Fig. 1(B for SEM and D for TEM) showed that large amounts of Au NPs as spherical particles about 10 nm were reduced on SiC NPs, indicating the successful preparation of Au@SiC nanocomposites. The size of Au@SiC NPs was determined by DLS using a Malvern Zetasizer Nano-ZS90. From the size distribution histogram of as shown in Fig. 1E, an average size of 370 nm for Au@SiC was obtained. Since the size of SiC and Au NPs demonstrated by TEM was 45 nm and 10 nm, respectively. It can be concluded that the 370 nm size for Au@SiC determined by DLS was the aggregation of several Au@SiC NPs. Fig. 1F showed the low-temperature nitrogen adsorption– desorption isotherms of SiC, which gave a clear IV type adsorptive isothermal curves, the adsorptive capacity rise sharply when the relative pressure is in the range from 0.80 to 0.95, and the hysteresis loop existed at the desorption process. As we known, the shapes of the isotherms and the hysteresis loop suggested the sample possesses a mesoporous structure. Additionally, above the relative pressure of 0.9, the isotherm still rise, indicating the sample also possesses some macro porous structures. The BET surface areas of the SiC was 40.0 m2/g, which was two times bigger than previously reported (Dai et al., 2012). The results illustrated that the SiC could be suitable for anchoring metal nanoparticles. FTIR spectra was employed to investigate carboxyl functionalization of SiC. Fig. 1G showed FTIR spectra of SiC and SiC–COOH. As observed, SiC displayed a strong band at 829 cm 1, which was ascribed to the typical stretching vibration of Si–C. The two weak bands at 3460 and 1640 cm 1 could be attributed to the stretching vibration and bending vibration of O–H from absorbed water. Compared with that of SiC, the FTIR spectrum of SiC–COOH displayed the presence of C ¼O (1724 cm 1), which confirmed the presence of a carboxyl group. The crystal structure of the Au@SiC nanocomposite was investigated by XRD. Fig. 1H showed the XRD patterns of SiC and Au@SiC nanocomposite. XRD patterns well defined peaks at 35.64, 41.34, 60.03, 71.82, and 75.33 1 (2θ), indicating the formation of the cubic phase of SiC NPs. According to the cubic phase crystal structure of SiC, we can conclude that the crystal form of the SiC is β-SiC (3C–SiC). XRD patterns well defined peaks at 44.35, 64.53, 77.57, and 81.86 o (2θ), indicating the formation of the cubic phase of Au NPs. The characteristic diffraction peak of Au NPs confirmed the presence of Au NPs in the Au@SiC nanocomposites.
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3.2. Electrochemical characterization of immunosensor To characterize the fabrication process of the immunosensor, CVs at all immobilization steps were recorded. As shown in Fig. S1, the Au@SiC–CS film modified GCE showed cathodic current at about –0.65 V in 0.1 M pH7.4 PBS. There is no peak current in the present sweep range from –0.3 V to 0.7 V. This may be attributed to the synergistic effect between Au NPs and SiC. As shown in Fig. 2A, a pair of typical reversible redox peaks of ferricyanide ions can be observed on the bare GCE (curve a) in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl. When a film of CS was immobilized onto a bare GCE, the current response increased obviously due to the fact that CS can facilitate electron transfer (Sassolas et al., 2012) between the solution and the electrode (curve b). When KL–CS, SiO2–CS, and SiC–CS were immobilized onto a bare GCE, SiC–CS/GCE (curve e) showed higher current response than KL–CS/GCE (curve c) and SiO2–CS/GCE (curve d), indicating that SiC is a good promoter of electron transport, and favored for electron communication between the solution and the electrode, which is consistent with previously described (Dai et al., 2012; Wright and Horsfall, 2007; Oliveros et al., 2013). Furthermore, the excellent adsorptive capacity of SiC can also lead to the higher current response. After Au NPs were reduced on SiC, the peak current response (curve f) increased drastically, improving the electron transfer from the redox probe to the electrode. The significantly enhanced current response could be attributed to the fact that Au NPs with splendid conductivity and large surface area could amplify the electrochemical signal (Wu et al., 2013). When anti-hCG was adsorbed onto the electrode through cross-link with the amino group of CS and carboxyl group of SiC, there was an obvious decrease of the current response (curves g), which suggested that antibody had been successfully immobilized on the electrode surface. Moreover, Au NPs with good biocompatibility
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were used to immobilize anti-hCG, Au NPs in the Au@SiC nanocomposites played a crucial role in antibody immobilization. The reason for the decrease of the current response is that anti-hCG as protein biomacromolecules insulated the conductive support and played a key role as obstructions in the process of electron transfer (Fan et al., 2013). While BSA was added onto the electrode, the current response was decreased further for the same reason (curves h). Additionally, the peak current (curve i) further decreased after incubation with 500 IU/L hCG, indicating the formed immunocomplex after specific immunoreaction acted as an electron-transfer and mass-transfer blocking layer, which greatly inhibited the electron transfer toward the electrode surface (Sun et al., 2013). From these results, it could be seen that the Au@SiC–CS film modified electrode could be used for the detection of hCG. The electrochemical behaviors of the fabrication process of the immunosensor were confirmed by DPV. The similar electrochemical characterizations were obtained as shown in Fig. 2B. The kinetics of the electrode reactions was investigated by studying the effect of scan rate at the BSA/anti-hCG/Au@SiC–CS immunoelectrode in 0.1 M pH 7.4 PBS containing 2 mM [Fe (CN)6]3 /4 and 0.1 M KCl. As shown in Fig. 3A, both anodic peak current (Ipa) and cathodic peak current (Ipc) increased with the increase of scan rate in the range of 10–100 mV/s. Also, the Ipa and Ipc showed a linear relationship with the square root of the scan rate respectively (Fig. 3B), suggesting that the electrode reaction is a diffusion-controlled electrochemical process, which is the ideal case for quantitative measurements (Mobin et al., 2010). It can be seen that both Ipa and Ipc of the immunoelectrode increased linearly and were proportional to the square root of the scan rate according to Eqs. Ipa (mA)¼ 9.05v1/2 (mV/s)1/2 þ 10.2 and Ipc (mA)¼10.4v1/2 (mV/s)1/2–13.9. The separation of peaks suggests that the process is not perfectly reversible; however, stable redox peak current and position during repeated scans at a particular
Fig. 3. CV studies of BSA/anti-hCG/Au@SiC–CS/GCE immunoelectrode as a function of scan rate (10–100 mV/s) (A); Magnitude of current response vs. square root of scan rate (B); Value of redox potential vs. scan rate (C); Value of redox potential vs. the Napierian logarithm of scan rate (D).
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scan rate suggests that the BSA/anti-hCG/Au@SiC–CS immunoelectrode exhibits a quasi-reversible process (Topoglidis et al., 2003). Moreover, both anodic peak potential (Epa) and cathodic peak potential (Epc) demonstrated a linear relationship with the scan rate respectively (Fig. 3C), indicating that the electron transport form redox moieties to the electrode is very facile (Kaushik et al., 2013). Epa and Epc were proportional to the scan rate according to equations. Epa (V)¼2.42v (mV/s) þ 0.206 and Epc (V) ¼ 2.87v (mV/s) þ 0.162. Fig. 3D showed the relationship between the peak potentials and the Napierian logarithm of scan rate (ln v). It can be observed that the scan rate affects the position of the redox peaks. With increasing the scan rate, the Epa shifted positively and the Epc shifted negatively. When v is small, both the anodic and cathodic peaks become closed to reversible peaks (Yin et al., 2010). However, the reversibility of electrode reaction changes with the scan rate. The peak separation (ΔEp) may enlarge as the scan rate becomes higher, which indicates the irreversibility of the electrode process increases. The fact that the ΔEp increases with increasing the scanning rate can be attributed to the polarization of the electrodes (Li et al., 2011b). Similarly, a linear relationship between peak potentials and Napierian logarithm of scan rate (ln v) is also observed in the scan rates ranging from 40 to 100 mV/s. The apparent heterogeneous electron transfer rate constant (ks) for the BSA/anti-hCG/Au@SiC–CS immunoelectrode was estimated using the Laviron model (Laviron, 1979) as ks ¼mnFv/RT, where m is peak-to-peak separation, F is Faraday constant, v is scan rate (mV/s), n is the number of transferred electrons and R is gas constant. The value of ks obtained as 3.9 s 1 (T ¼298 K, n ¼1, m ¼0.096 V and v ¼100 mV) is higher than that of other nanoparticles based bioelectrodes (Zhang et al., 2005; Zhao et al., 2005), indicating fast electron transfer between immobilized biomolecules and electrode. 3.3. Optimization of experimental conditions To achieve the optimal immunoassay performance, the pH of the detection solution, the immunoreaction temperature, and immunoreaction time were optimized as important factors that influenced the sensitivity of the proposed immunosensor. The optimization of experimental conditions was supplied in Supplementary Materials. 3.4. Calibration curve of immunosensor
Fig. 4. DPVs of the immunosensor incubated with different concentrations of hCG standard solution (from a to j): 0.0, 1.0, 10.0, 50.0, 100.0, 200.0, 300.0, 500.0, 700.0, and 1000.0 IU/L in 0.1 M pH 7.4 PBS containing 2 mM [Fe(CN)6]3 /4 and 0.1 M KCl. Pulse width: 0.05 s; amplitude: 0.05 V (A); Calibration plots of the reduction current change (ΔI) versus concentration of hCG under optimal conditions (B); Selectivity of the proposed immunosensor with FSH (20 IU/L), LH (20 IU/L), TSH (20 mIU/L), hCG (5 IU/L), and hCG (5 IU/L) containing the above mixture of three interferents with the same concentrations (C).
Under the optimal detection condition, the immunosensor was subjected to standard hCG solution with various concentrations. As indicated in Fig. 4A, the peak currents decreased with the increase of hCG concentration, which may be ascribed to that more antigen– antibody immunocomplex as an insulating layer was formed on the surface of electrode, and thus inhibiting the electron transfer. Therefore, the current change was proportional to the concentration of the corresponding hCG. The corresponding calibration curve was plotted in Fig. 4B. The current change was linearly related to hCG concentration in the range of 0.1–5 IU/L and 5–1000 IU/L. Correspondingly, the regression equations were ΔI (mA)¼0.81C (IU/L)þ 2.05 and ΔI (mA)¼ 0.036C (IU/L)þ6.11 with correlation coefficients of 0.996 and 0.998. From the slope of 0.81 the detection limit was calculated to be 0.042 IU/L (signal/noise [S/N]¼3). Compared with other sensors reported previously (Chai et al., 2008; Wang et al., 2010; Yang et al., 2011a; 2011b), the proposed immunosensor exhibited a satisfactory detection limit and linear range. The characteristics of other hCG sensors are summarized in Table 1. Lu et al. reported a sandwich-type immunosensor based on Au-multiwalled carbon nanotubes-graphene composite modified GCE with a very low detection limit of 0.0026 IU/L. However, the fabrication process
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Table 1 Comparison of different immunosensors for detection of hCG. Electrode
Method
Liner range (IU/L)
Detection limit (IU/L)
Ref
anti-hCG/nano-Au/MBa/GCE anti-hCG/nano-gold and CS hybrid film/GCE anti-hCG/Pt–Au alloy nanotube array/GCE HRPb-anti-hCG-hCG modified electrode HRP-anti-hCG/hCG/SGc/GCE Gold-labeled-anti-hCG-Ag(I)-hydroquinone HRP-anti-hCG /hCG/AuNPs/MPSd/GCE HRP-anti-hCG/sol–gel/GEe anti-hCG/GNPsf/pPAg/MWCNTsh/GCE anti-hCG/gold nanotubes array/GCE anti-hCG/Au/MWCNTs/GSi/GCE anti-hCG/AuNPs-TiO2/Thij/GAk/MWCNTs-CS/GCE anti-hCG/Au@SiC-CS/GCE
CV CAl CAl DPV DPV RS assayl DPV DPV CV CAm DPV DPV DPV
1–100 0.2–100 25–400 2.5–12.5 0.5–5.0; 5.0–30 2.5–208.3 5.0–30 0.5–50 1–10, 10–160 0.1–100 0.005–500 0.2–300 0.1–5, 5–1000
0.3 0.1 12 1.4 0.3 0.83 1.4 0.3 0.3 0.08 0.0026 0.08 0.042
Chai et al. (2008) Yang et al. (2009) Tao et al. (2011) Chen et al. (2005) Chen et al. (2006a) Liang et al. (2008) Chen et al. (2006b) Tan et al. (2007) Wang et al. (2010) Yang et al. (2011b) Lu et al. (2012) Yang et al. (2011a) This work
a
Methylene blue. Horseradish peroxidase. c Titania sol–gel matrix. d 3-Mercaptopropanesulfonic acid. e Graphite electrode. f Gold nanoparticles. g Poly-(2,6-pyridinediamine). h Multiwalled carbon nanotubes. i Graphene nanosheets. j Thionine. k Glutaraldehyde. l Chronoamperometry. m Resonance scattering assay. b
increased the complexity of the operation system. In the present work, such a high sensitivity of the proposed immunosensor may be attributed to three factors: (1) SiC’s good biocompatibility, electrical, mechanical, and thermal properties determine its suitability as a biomaterial and biosensing substrate (Wright and Horsfall, 2007); (2) Au NPs possess splendid conductivity, large surface area, and good biocompatibility, which could amplify the electrochemical signal, resulting in high sensitivity, and enhance the stability of the immunosensor; (3) the large amounts of antibodies have been immobilized onto the electrode by Au NPs, the amino group of CS, and carboxyl group of SiC. Also, CS produced a thin film with good biocompatibility and high chemical stability, which increased the sensitivity of the immunosensor.
3.5. Selectivity, reproducibility, and stability It has been reported that some potential interferents such as carcinoembryonic antigen, α-1-fetoprotein, prostate-specific antigen, cancer antigen 125, carbohydrate antigen 19-9, ferritin, BSA, ascorbic acid, L-cysteine, and L-glutamic acid have no interference to the detection of hCG (Wang et al., 2010; Yang et al., 2011a; Tan et al., 2007; Lu et al., 2012; Wu et al., 2013). Thus, three sister molecules of hCG, namely, FSH (20 IU/L), LH (20 IU/L), and TSH (20 mIU/L), which are the known interferents in the hCG detection system, were selected as the predominant interferents in this study. Selectivity was investigated by recording current changes of the proposed immunosensor to the three potential interferents. The result was expressed in Fig. 4C. Compared to the result obtained from the only hCG (5 IU/L), the change of current response of the immunosensor before and after incubation with FSH, LH, and TSH could be neglected. These demonstrated that the three interferents did not cause observable interference to hCG detection, which was attributed to the highly specific antigen–antibody immunoreactions. Additionally, the cross-selectivity of the immunosensor incubated with mixture consisting of hCG and three interferents was also investigated. No remarkable change of ΔI obtained from mixture was observed in comparison with that in the presence of only hCG.
These results indicated that the proposed immunosensor specifically recognized hCG antigen and exhibited good selectivity. The reproducibility of the immunosensor was examined by six equally proposed immunosensors incubated with the same concentration hCG (100 IU/L). The six electrodes exhibited the similar electrochemical responses and a relative standard deviation (RSD) of 3.8% was obtained, indicating satisfying reproducibility. Successive cyclic potential scans for 50 cycles and long-term storage assay were used to examine the stability of the proposed immunosensor. A 4.4% decrease of initial peak current was found after 50 continuous cycle scans. Additionally, the long-term stability experiment was carried out intermittently (every 5 days). When the immunosensor was not in use, it was stored in a refrigerator at 4 1C. Over 95.6% and 88.5% of initial response remained after storage of 15 and 30 days, respectively. The acceptable stability of the immunosensor may be due to the facts that Au NPs loaded on the electrode surface could make protein molecules more firmly attached on electrode. CS as dispersant of Au@SiC nanocomposite had outstanding film-forming ability and high chemical stability. 3.6. Real sample analysis In order to evaluate the feasibility of the proposed immunosenor for real sample analysis, the immunosensor was used for the determination of hCG by standard addition methods in serum samples. The result showed that the recovery was from 98.7% to 103.6% and the RSD was from 1.7% to 7.0% (Table S1), which confirmed the practical value of the immunosensor. In the present work, a simple and effective label-free electrochemical immunosensor based on Au@SiC–CS film was developed. The proposed immunosensor provides a simple and economic method with higher sensitivity and a wider concentration range for detection of hCG. Thus, we believe that there are many opportunities for SiC as an active material to be incorporated in such devices that require certain chemical, mechanical, and electrical properties, which could find potential application in clinical analysis. In addition, the bio-functionalization of SiC
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combined with other nanocomposite and the ongoing research toward biosensor configurations that employ SiC as an active material will increase the possibilities to create complex devices that can perform multiple biomolecule detection and analysis on a single platform. Thus, great care must be taken when profiling multiple analytes simultaneously. 4. Conclusion In summary, the uniform and highly dispersed Au@SiC nanocomposites were synthesized by a facile ascorbic acid assisted chemical reduction. A simple and effective label-free electrochemical immunosensor based on Au@SiC–CS film for hCG detection was developed. The immunosensor displayed a linear response for detection hCG within a wide range (0.1–1000 IU/L). The proposed immunosensor shows a low detection limit (0.042 IU/L), good reproducibility, selectivity, and acceptable stability. The simple fabrication procedure and the ultrasensitivity demonstrated by the immunosensor may provide many potential applications for the detection of hCG in clinical diagnostics. Acknowledgments This work was supported by the Natural Science Foundation of China (31160334; 31260226), the Natural Science Foundation of Yunnan Province (2012FB112; 2012FB121), and the Scientific Research Fund of the Yunnan Provincial Education Department (2011Z051) People’s Republic of China. Appendix A. Supplementary material Supplementary data associated with this article can be found in the online version at http://dx.doi.org/10.1016/j.bios.2014.02.019. References An, Y.Q., Chen, M., Xue, Q.J., Liu, W.M., 2007. J. Colloid Interf. Sci. 311, 507–513. Belmonte, M., Nieto, M.I., Osendi, M.I., Miranzo, P., 2006. J. Eur. Ceram. Soc. 26, 1273–1279. Chai, R., Yuan, R., Chai, Y.Q., Ou, C.F., Cao, S.R., Li, X.L., 2008. Talanta 74, 1330–1336. Chen, J., Yan, F., Dai, Z., Ju, H.X., 2005. Biosens. Bioelectron. 21, 330–336. Chen, J., Tang, J.H., Yan, F., Ju, H.X., 2006a. Biomaterials 27, 2313–2321. Chen, J., Yan, F., Tan, F., Ju, H.X., 2006b. Electroanalysis 18, 1696–1702. Chu, V., Conde, J.P., Jarego, J., Brogueira, P., Rodriguez, J., Barradas, N., et al., 1995. J. Appl. Phys. 78, 3164–3173. Cui, R.J., Pan, H.C., Zhu, J.J., Chen, H.Y., 2007. Anal. Chem. 79, 8494–8501. Dai, H., Chen, Y.L., Lina, Y.Y., Xu, G.F., Yang, C.P., Tong, Y.J., Guo, L.H., Chen, G.N., 2012. Electrochim. Acta 85, 644–649. Deva Reddy, J., Volinsky, A.A., Frewin, C.L., Locke, C., Saddow, S.E., 2008. In Materials Research Society Meeting, ed. Mechanical properties of single and polycrystalline SiC thin films, pp. AA03–AA6. Fan, H.X., Zhang, Y., Wu, D., Ma, H.M., Li, X.J., Li, Y., Wang, H., Li, H., Du, B., Wei, Q., 2013. Anal. Chim. Acta 770, 62–67. Ferroni, L.P., Pezzotti, G., 2002. J. Am. Ceram. Soc. 85, 2033–2038. Huang, K.J., Li, J., Wu, Y.Y., Liu, Y.M., 2013. Bioelectrochemistry 90, 18–23. Kaushik, A., Vasudev, A., Arya, S.K., Bhansali, S., 2013. Biosens. Bioelectron. 50, 35–41.
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