toluidine blue–graphene oxide nanocomposites

Sensors and Actuators B 207 (2015) 269–276

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Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb

Label-free electrochemical DNA biosensor for rapid detection of mutidrug resistance gene based on Au nanoparticles/toluidine blue–graphene oxide nanocomposites Hua-Ping Peng a,c , Yan Hu a,c , Pan Liu a,c , Ya-Ni Deng a,c , Peng Wang a,c , Wei Chen a,c , Ai-Lin Liu a,c,∗ , Yuan-Zhong Chen b,∗∗ , Xin-Hua Lin a,c,∗ a

Department of Pharmaceutical Analysis, Faculty of Pharmacy, Fujian Medical University, Fuzhou 350004, China Fujian Key Lab of Hematology, Fujian Institute of Hematology, Affiliated Union Hospital of Fujian Medical University, Fuzhou 350000, China c Nano Medical Technology Research Institute, Fujian Medical University, Fuzhou 350004, China b

a r t i c l e

i n f o

Article history: Received 11 July 2014 Received in revised form 3 October 2014 Accepted 14 October 2014 Available online 23 October 2014 Keywords: Label-free electrochemical DNA biosensor Multidrug resistance gene Au nanoparticles/toluidine blue–graphene oxide nanocomposites

a b s t r a c t The occurrence of multidrug resistance (MDR) has become a major obstacle to the successful performance of chemotherapy for cancer patients, so it is highly required to develop methods to explore the new strategy to early diagnose the MDR. Here, we report a novel label-free electrochemical DNA biosensor for simple, effective and convenient determination of MDR1 gene based on Au nanoparticles/toluidine blue–graphene oxide (Au NPs/TB–GO) modified electrode. The resulting Au NPs/TB–GO nanocomposites were characterized by scanning electron microscopy, atomic force microscope, ultraviolet–visible spectrometry, cyclic voltammetry, and electrochemical impedance spectroscopy. Differential pulse voltammetry was employed to monitor the hybridization of DNA by measuring the changes in the peak currents of TB. Under optimal conditions, the decreased currents were proportional to the logarithm of the concentration of the target DNA in the range of 1.0 × 10−11 –1.0 × 10−9 M with a detection limit of 2.95 × 10−12 M (at an S/N of 3). In addition, the biosensor exhibited good selectivity, acceptable stability and reproducibility. The proposed method was simple, fast and inexpensive for the determination of MDR1 gene at low levels. © 2014 Elsevier B.V. All rights reserved.

1. Introduction Cancer is one of the most serious threats worldwide to human health. Although chemotherapy has been one of most effective cancer treatments, the resistance of tumor cells to chemotherapeutic drugs is becoming a major obstacle [1]. In human cells, expression of the (multidrug resistance) MDR1 gene, encoding a transmembrane efflux pump (P-glycoprotein), leads to decreased intracellular accumulation and resistance to a variety of lipophilic drugs (multidrug resistance, MDR). The levels of MDR in cell lines selected in vitro have been shown to correlate with the steadystate levels of MDR1 mRNA and P-glycoprotein [2]. A variety of

∗ Corresponding authors at: Department of Pharmaceutical Analysis, Faculty of Pharmacy, Fujian Medical University, Fuzhou 350004, China. Tel.: +86 591 22862016; fax: +86 591 22862016. ∗∗ Corresponding author. E-mail addresses: [email protected] (A.-L. Liu), [email protected] (Y.-Z. Chen), [email protected] (X.-H. Lin). http://dx.doi.org/10.1016/j.snb.2014.10.059 0925-4005/© 2014 Elsevier B.V. All rights reserved.

techniques have been developed for the detection and diagnosis of the MDR in cancer, including northern blot, western blot, immunocytochemistry, flow cytometry, and reverse transcriptase polymerase chain reaction (RT-PCR), and functional MDR assay [2–5]. However, there were some limitations in these methods, such as time consuming, poor precision, technical difficulty and expensiveness [5]. Thus, it is urgently needed to develop an effective diagnostic protocol for MDR with good sensitivity, specificity, simplicity and inexpensiveness is very significant in clinical assay. As an alternative to the conventional procedures, electrochemical DNA biosensors have received a particular attention because of its low cost, fast response, good selectivity, portability and simple instrumentation [6–8]. To date, various electrochemical DNA biosensors have been reported. To our knowledge, developing of label-free electrochemical DNA biosensors has been proved to be one successful method to directly detect target DNA [9–11]. As is well known, toluidine blue (TB) is a kind of phenothiazine dye and has good reversible electron transfer ability. TB has been widely used as a feasible redox probe for electrochemical sensors due to its good chemical stability and high electrical conductivity

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[12,13]. However, leakage is a main problem in the entrapment of mediator (low molecular weight) compounds in the matrices. This problem can be effective solved by covalent linking of mediator with polymers, or with high molecular weight compounds before immobilization. Especially, covalent linking of mediator with nanomaterials has greatly improved the sensitivity and selectivity of electrochemical biosensors [12,14]. Thus, a novel functional TBnanomaterial should be designed for effectively preventing the leakage of the mediator and improving the stability of the biosensor. In recent years, graphene is a novel, one-atom thick, twodimensional graphitic form of carbon that has drawn intense attention in graphene-based sensor by various sensing technologies because of its unique structure and easy conjugation to proteins without degrading their biological activity [15–20]. Graphene oxide (GO) sheets, as the water-soluble derivative of graphene, has been of increasing interest for applications in variety fields due to its unique characteristics, such as excellent dispersibility, good biocompatibility and facile surface functionality [21,22]. Compared with pristine graphene, the covalent oxygenated functional groups in GO can indeed give rise to remarkable structure defects. This is concomitant with some loss in electrical conductivity, which possibly limits the direct application of GO in electrically active materials and devices. However, in recent years, developing novel GO-based nanocomposite materials has been of increasing interest for applications in electrochemical biosensing. Due to its favorable electron mobility and unique surface properties, such as one-atom thickness and high specific surface area, GO can accommodate the active species and facilitate their electron transfer at electrode surfaces [23]. On the other hand, Au nanoparticles

(Au NPs) have also drawn much attention in electrochemical field because of its excellent characteristics, such as large surface area, high chemical stability, good biocompatibility, and being able to facilitate electron transfer between biomolecules and electrodes [24,25]. Moreover, Au NPs can link with a lot of compound through Au NH or Au S chemical bond or electrostatic interaction because of its negatively charge [26–28]. For these reasons, Au NPs have been widely used for fabrication of biosensors. In view of the advantageous features of TB, GO and Au NPs, herein we developed a facile and sensitive label-free electrochemical DNA biosensor interface based on Au NPs/TB–GO nanocomposites film for the effective detection of MDR1 gene. Au NPs/TB–GO nanocomposite films have the advantages of each element and particularly the following ones: (1) TB not only acted as an electrochemical active mediator due to its good electrochemical redox active properties, but also acted as a “glue” to connect the Au NPs to the GO sheets due to the interaction between the amine groups in TB and Au NPs together with the electrostatic interaction between TB and Au NPs; (2) GO was not only used to increase the loading capacity of the TB due to its large specific surface area, but also used to avoid the leakage of the TB from the electrode surface by forming the TB–GO nanocomposite; (3) Au NPs were used to immobilize the probe DNA via Au S bond. In addition, Au NPs could provide a favorable microenvironment to retain the bioactivity of the immobilized DNA and effectively promoted electronic transfer due to their excellent biocompatibility and the good conductivity. When the target DNA was hybridized with probe to form the double-stranded DNA (dsDNA), the electrochemical signal of TB decreased. The approach for the fabrication of the biosensor and detection of DNA was illustrated in Scheme 1.

Scheme 1. Fabrication and detection process of the DNA biosensor.

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2. Experimental 2.1. Materials Graphite flakes (99.99%, 325 mesh) was purchased from Alfa Aesar. Toluidine blue (TB) and 6-mercapto-1-hexanol (MCH) were purchased from sigma. Chloroauric acid (HAuCl4 ·4H2 O), trisodium citrate and tannic acid were obtained from Sinopharm Chemical Reagent Co. Ltd. (Shanghai, China). All oligonucleotides were synthesized and purified by TaKaRa Inc. (Dalian, China), and the sequences are (probe, P1 ) 5 -SH-TTC CTT CTT ATC TTT TTC ACT TTT ATT GTT-3 , (target, C1 ) 5 -AAC AAT AAA AGT GAA AAA GAT AAG AAG GAA-3 , (non-complementary, C2 ) 5 -AAC AAT AAA AGT GAA AGA GAT AAG AAG GAA-3 , (One-base mismatch, C3 ) 5 -CGA CCG TGC CTC AGC CTG CTA TCA CTA CCG-3 , (20-base mismatch, C4 ) 5 -AGG ACC ACC GCA GAT CTA CAT TCA AGA ACT-3 . The buffer solutions used in this study were as follows: the DNA immobilization buffer contained 10 mM Tris–HCl, 1 mM ethylenediaminetetraacetic acid, and 1 M NaCl (pH 8.0). The hybridization buffer was a 10 mM phosphate buffer solution (PBS) containing 1 M NaCl (pH 7.0). All solutions were prepared with Milli-Q water (18 M cm resistivity) from a Millipore system. 2.2. Preparation of TB–GO NC Water-soluble GO was prepared from graphite flakes following the method described by a modification of Hummers [29–31]. The TB–GO NC was prepared as follows. The GO powder was treated by ultrasonication for 1 h then centrifuged at 10,000 rpm for 10 min to remove the supernatant to obtain brown GO aqueous. Subsequently, TB–GO NC was prepared by using sonochemical approach. In brief, 10 mg mL−1 of TB aqueous was added dropwise to the 500 ␮L of GO aqueous (2 mg mL−1 ) while ultrasonicating vigorously to avoid aggregating. The color of TB–GO NC was light green at first, and then turned to blue. TB–GO NC was then obtained through water washing and centrifugation for several times, and the resulting TB–GO NC was suspended in 1 mL PBS (pH 7.0) and stored at 4 ◦ C when not used. 2.3. Preparation of Au NPs Au NPs were synthesized via reduction of HAuCl4 ·4H2 O by trisodium citrate and tannic acid [32,33]. The ratio of tannic acid to trisodium citrate determines the size and distribution of synthesized Au NPs. Two initial aqueous solutions (a) 1 mL of 2 wt.% HAuCl4 ·4H2 O solution, which was diluted to 80 mL, and (b) a mixture of 8 mL of 1 wt.% trisodium citrate and 0.2 mL of 1% tannic acid, which was diluted to 20 mL, were heated to 60 ◦ C on a hot plate. Then, solution (b) was rapidly added to solution (a) under rigorous stirring for 35 min to give Au NPs solution. The concentration of as-synthesized Au NPs was determined to be 0.12 mg mL−1 on the basis of the concentration of HAuCl4 ·4H2 O before reaction, assuming a 100% reduction yield of the [AuCl4 ] ions [34]. 2.4. Fabrication of DNA electrochemical biosensor Glassy carbon electrode (GCE, 3 mm diameter, CH Instruments, Inc.) was wet polished down to mirror-like with 1.0, 0.3 and 0.05 ␮m alumina slurry, respectively, and rinsed thoroughly with doubly distilled water between each polishing step. The electrode was then successively sonicated in 1:1 nitric acid, acetone and doubly distilled water, and then allowed to dry at room temperature. After mixing 20 ␮L of TB–GO suspension with 3 ␮L of 1.5% nafion solution, 5 ␮L of the mixture was uniformly cast onto the well-polished GCE surface, and then drying slowly in air at room temperature. Subsequently, the obtained electrode was immersed

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into the Au NPs solution for 4 h and the Au NPs were absorbed on TB–GO surface, obtaining an Au NPs/TB–GO nanocomposite. After washing with water, a 5 ␮L volume of 1 × 10−6 M probe DNA solution was drop cast to cover the modified gold electrode and incubated for 2 h at room temperature. Probe DNA with thiol groups at 5-end were covalently linked onto the surface of the Au NPs/TB–GO modified electrode via Au–S binding. The singlestranded DNA (ssDNA) modified electrode was then treated with 1.0 mM MCH for 1 h, followed by thoroughly washing with deionized water to remove unspecific adsorbed DNA probe. The modified electrode was donated as P1 /Au NPs/TB–GO/GCE. 2.5. Hybridization and electrochemical measurements of the biosensor The hybridization procedure was performed by immersing the probe DNA modified electrode into 0.01 M PBS containing various concentrations of target DNA for 1 h at 33 ◦ C, and then the hybridized electrode was rinsed with PBS to remove the nonspecifically adsorbed DNA. The obtained electrode after hybridization with target DNA was denoted as C1 –P1 /Au NPs/TB–GO/GCE, C2 –P1 /Au NPs/TB–GO/GCE, and C3 –P1 /Au NPs/TB–GO/GCE and used for electrochemical measurements. The fabrication and detection process of the DNA biosensor was schematically illustrated in Scheme 1. 2.6. Characterization The electrochemical measurements for electrochemical impedance spectroscopy (EIS), cyclic voltammogram (CV) and differential pulse voltammetry (DPV) were performed on an Autolab PGSTAT30 electrochemical workstation (Eco Chemie). A three-electrode system was used with an Ag/AgCl electrode as the reference electrode, a platinum wire as the auxiliary electrode, and the modified glassy carbon electrode (GCE) as the working electrode. The electrochemical impedance spectroscopy (EIS) measurements were performed in the solution of 0.10 M KCl containing 5 mM [Fe(CN)6 ]4−/3− as a redox probe at an open potential of 210 mV within the frequency range of 10−2 to 105 Hz. The amplitude of the applied sine wave potential was 5 mV. Differential pulse voltammetry (DPV) was carried out from −0.50 to 0.05 V with pulse amplitude of 50 mV. Atomic force microscope (AFM) image was carried out on Bruker NanoScope V (Germany). The AFM sample was prepared by casting the samples onto a freshly cleaved mica surface, followed by drying at room temperature. UV–vis absorption spectra were acquired with a UV 2450 spectrophotometer (Shimadzu, Japan). 3. Results and discussion 3.1. Characterization of the Au NPs/TB–GO nanocomposites The morphologies of GO, TB–GO and Au NPs/TB–GO films were characterized by SEM and AFM techniques (Fig. 1). As can be seen from Fig. 1a, the SEM image and AFM image of GO showed a planar sheet-like structure, indicating that GO had been readily exfoliated into individual sheets in water by ultrasonic treatment. The height of GO was around 0.8 nm which coincided with the literature was typical for single-layer sheet [35]. When TB was modified on the surface of GO (Fig. 1b), as expected, the TB–GO hybrid film displayed a rough surface structure. After absorption of Au NPs, many pearl-like particles with diameter of ∼30 nm distributed uniformly onto the TB–GO matrix (Fig. 1c), indicating the strong interaction between the Au NPs and the TB–GO nanocomposite [36]. UV–vis spectroscopy was conducted to evaluate the formation of Au NPs/TB–GO NCs. As shown in Fig. 2, the GO solution showed

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Fig. 2. UV–vis absorption spectra of GO (a), TB (b), TB–GO (c), Au NPs/TB–GO (d) and Au NPs (e) solution.

a characteristic band at 230 nm and a shoulder at 300 nm (curve a), which corresponded to the ␲–␲* transition and n–␲* transition, respectively [37]. The spectrum of TB–GO NCs solution exhibited two absorption peaks around 287 nm and 584 nm (trace c). In comparison with those obtained at pure TB solution (trace b), the blue shift and broadened for the absorption at 584 nm was observed. This phenomenon could be mainly due to the electrostatic interactions between GO and TB and the increased steric hindrance for the free conformation of TB on GO after TB molecules adsorbed onto GO sheets. Besides, the spectrum of Au NPs solution showed a characteristic band at 520 nm (trace e). When the Au NPs absorbed onto the surface of TB–GO nanocomposite, the absorption peak of Au NPs shifted to ∼550 nm and became a broader absorption band. The changes in the optical properties of the Au NPs were consistent with other previous studies [38]. In addition of the characteristic band of TB, a wide absorption peak at 550–610 nm was achieved (trace d), which might be ascribed to the absorption both of the nanosized Au NPs and TB in Au NPs/TB–GO NCs [39]. Thus, the UV–vis absorbance spectrum confirmed the immobilization of the TB and the Au NPs on the surface of GO nanosheets. 3.2. Electrochemical characterization of the different modified electrodes

Fig. 1. SEM images of GO (a), TB–GO (b), and Au NPs/TB–GO (c). Inset of (a) and (c): AFM images of GO and Au NPs/TB–GO nanocomposite.

Cyclic voltammetry (CV) is an effective and convenient method for probing the feature of the modified electrode surface. The CVs of different modified electrodes in the potential range from −0.5 to 0.05 V in 0.1 M PBS (pH 7.0) were shown in Fig. 3. No obvious redox peaks were observed at the bare GC electrode (curve a). When TB–GO NC was immobilized onto the surface of modified electrode, a couple of well-defined redox peaks at −296 and −225 mV were obtained (curve b). This is in accordance with the literature [40], revealing that the TB attached on GO surface retained the redox properties. Absorption of Au NPs increased the conductivity of the TB–GO film, resulting in the increase of the redox peak current as the Au NPs acted as a conducting wire or an electron communication relay, which increased the electron transfer efficiency (curve c) [27,41]. When probe DNA was immobilized on the surface of the electrode, the redox peak current of TB decreased. After hybridization with target DNA, the peak current of the modified electrode further decreased, indicating that dsDNA formed an electron-transfer and mass-transfer blocking layer, which greatly inhibited the reaction on the electrode surface [27,42]. The CVs of

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Fig. 4. Nyquist plots of the different electrodes in a PBS (0.1 M, pH 7.0) solution containing 0.1 M KCl and 10 mM Fe(CN)6 3−/4− : (a) bare GCE, (b) TB–GO/GCE, (c) Au NPs/TB–GO/GCE, (d) P1 /Au NPs/TB–GO/GCE, (e) C1 –P1 /Au NPs/TB–GO/GCE. Rs , Zw , Ret and Cdl represent the solution resistance, the Warburg diffusion resistance, the electron transfer resistance and the double layer capacitance, respectively.

Fig. 3. (A) CVs of different modified electrodes in pH 7.0 PBS at a scan rate of 100 mV s−1 . (a) Bare GCE, (b) TB–GO/GCE, (c) Au NPs/TB–GO/GCE, (d) P1 /Au NPs/TB–GO/GCE, (e) C1 –P1 /Au NPs/TB–GO/GCE. (B) CVs of the electrode (e) in 0.1 M pH 7.0 PBS at 10, 20, 50, 70, 100, 120, 150, 170, 200 and 250 mV s−1 (from inner to outer). Inset: Plots of peak current vs. scan rate.

the DNA biosensor in 0.1 M PBS (pH 7.0) at different scan rates were also studied. The results were shown in Fig. 3B. Obviously, the couple of redox peak currents increased with the increase of potential scan rate. In addition, both the anodic and cathodic peak currents were directly proportional to the potential scan rates in the range of 10–250 mV s−1 (shown in the inset), suggesting a surface confined redox process [43]. 3.3. EIS characterization of the DNA biosensor It is well known that EIS is an excellent technique for investigating the interface properties of surface-modified electrodes [44,45]. A typical impedance spectrum (presented in the form of the Nyquist plot) exhibits a semicircle near the origin at higher frequencies corresponding to the electron transfer-limited process and a linear tail at lower frequency range representing the diffusion-limited process. The electron-transfer resistance, Ret , is the most directive and sensitive parameter that responds to changes on the electrode interface, as represented by the diameter of the semicircle in the Nyquist plot. Fig. 4 showed the impedance spectra corresponding to the stepwise modification processes. The data could be fitted with a modified Randles equivalent circuit (inset in Fig. 4). A small

Ret was observed at the bare GCE (Fig. 4a). After nafion/TB–GO film was modified on the electrode, the Ret increased to 940  (Fig. 4b). When Au NPs were assembled on the modified electrode, the Ret decreased to 590  (Fig. 4c), this may be ascribed to that Au NPs acted as high electron relay for shuttling electron between the electrochemical probe and the electrode. The Ret increased when the DNA probe was immobilized, this was because the negatively charged phosphate backbone of the probe DNA immobilized on the electrode repelled the [Fe(CN)6 ]3−/4− anions [46]. Subsequently, the target DNA was hybridized with probe to form the dsDNA, and the value of Ret increased to 2093  due to large amount of negatively charged DNA linked on the electrode surface (Fig. 4e). These above results indicated that the DNA sensor has been successfully prepared and DNA hybridization could happen on the electrode surface.

3.4. Optimization of assay conditions As it is well known that the assay conditions affect the hybridization efficiency of a DNA biosensor. In this work, we selected the decline of the peak current of the probe modified electrode before and after hybridization with target DNA (I) as the response signal to investigate the effect of the experiment’s conditions on the performance of the biosensor. The effect of the hybridization temperature on the response signal was firstly investigated. As shown in Fig. 5A, a series of hybridization temperatures in a range of 28–45 ◦ C were evaluated. The maximum response signal was observed at around 33 ◦ C. When the hybridization temperature increased continuously, the response signal decreased quickly. Therefore, 33 ◦ C was used for hybridization of the probe and target in all following experiments. The effect of DNA hybridization time on the response signal was also investigated (Fig. 5B). It was observed that the response signal increased significantly as the hybridization time increased from 30 to 60 min, and remained constant after 60 min, indicating that the hybridization reaction was to a great extent completed after 60 min. From the perspective of sensitivity and assay time, we selected 60 min as the optimum DNA hybridization time.

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Fig. 6. DPV of P1 /Au NPs/TB–GO/GCE (a), C2 –P1 /Au NPs/TB–GO/GCE (b), C3 –P1 /Au NPs/TB–GO/GCE (c) and C1 –P1 /Au NPs/TB–GO/GCE (e) in 0.1 M PBS (pH 7.0). All the concentrations of different hybridized sequences were 5.0 × 10−10 M.

detection limit of 2.95 × 10−12 M (at an S/N of 3). The linear range and detection limit of this biosensor was compared with that published in the similar work previously (Table 1), showing that the proposed DNA sensor has good analytical performances for the specific sequences of DNA detection. 3.7. Reproducibility and stability of the DNA sensor

Fig. 5. Effects of (A) incubation temperature, and (B) incubation time on the DNA biosensor.

3.5. Selectivity of DNA biosensor Selectivity is a crucial factor to evaluate the performance of a DNA biosensor. In this work, the selectivity was explored using the P1 /Au NPs/TB–GO/GCE to hybridize with different sorts of DNA sequences. Fig. 6 shows the DPV measured following the hybridization of different kinds of target ssDNA (1.0 × 10−10 M) with probe modified electrode. After hybridization with non-complementary ssDNA (C2 ) and single-base mismatched ssDNA (C3 ), I were −0.043 ␮A and −0.96 ␮A, respectively, which correspond to 1.98% and 44.23% of the signal for complementary target ssDNA (C1 , −2.17 ␮A). These results showed that the fabricated DNA biosensor performed good selectivity.

To evaluate the reproducibility of the DNA sensor, a series of five electrodes were prepared for detecting 5.0 × 10−9 M C1 sequence. The relative standard deviation (RSD) of the measurements for the five electrodes was 4.6%, which shows the high reproducibility of the DNA detection. The stability of the successive assays was studied. After 50 CV measurements in PBS, a RSD of 3.3% was acquired. The long-time stability of the DNA sensor was also investigated. When not in use, the electrode was stored in PBS at 4 ◦ C. After two weeks, the response of the DNA sensor retained about 90% of its initial value. The good stability might be due to the fact that the DNA sequences were attached firmly onto the surface of Au NPs/TB–GO nanocomposites film. 3.8. Analytical application of the biosensor Another attractive feature of the developed method is that it can be used for assaying the MDR1 gene level precisely from the mixture of C1 , C2 , C3 , and C4 . Three samples containing a mixture of

3.6. Analytical performance Under the optimal conditions, the analytical performance of the DNA biosensor was investigated using the probe DNA to hybridize with the different concentrations of target DNA sequences. Fig. 7 shows the DPVs of the probe modified electrode at various complementary target DNA concentrations. As expected, the peak current decreased as the concentrations of the complementary target ssDNA increased. The decrease of peak current (I) was linear with the logarithm of the concentration of the complementary target ssDNA in the range from 1.0 × 10−11 M to 1.0 × 10−9 M, with a

Fig. 7. DPV of C1 –P1 /Au NPs/TB–GO/GCE varying the concentration of C1 from 0 M (a), 1.0 × 10−11 M (b), 5.0 × 10−11 M (c), 1.0 × 10−10 M (d), 5.0 × 10−10 (e), and 1.0 × 10−9 (f) in 0.1 M PBS (pH 7.0). Inset: the calibration curve of the biosensor.

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Table 1 Comparison of linear ranges and detection limits of the different electrochemical DNA sensors. Modified electrodes

Detection technique

Linear range (M)

Detection limit (M)

References

CNTs PNA/poly(JUG-co-JUGA) Nanoporous alumina lambda exonuclease MWCNTs-COOH Au NPs/TB–GO

DPV SWV DPV EIS DPV DPV

6.7 × 10−10 –8.4 × 10−9 1 × 10−8 –1 × 10−7 1.0 × 10−11 –1.7 × 10−6 1.0 × 10−10 –2.0 × 10−8 1.6 × 10−9 –4.8 × 10−8 1.0 × 10−11 –1.0 × 10−9

1.4 × 10−10 1 × 10−10 9.55 × 10−12 9.8 × 10−11 3.8 × 10−11 2.9 × 10−12

[47] [48] [49] [50] [51] This work

Table 2 Measurements of MDR1 gene level in the mixtures of C1 , C2 , C3 , and C4 using the developed method. Sample no.

Added amounts of MDR1 gene (M)

Measured amounts of MDR1 gene (M)a

Recovery (%)

1 2 3

1.0 × 10−9 1.0 × 10−10 1.0 × 10−11

(0.94 ± 0.05) × 10−9 (1.03 ± 0.07) × 10−10 (1.08 ± 0.12) × 10−11

94 103 108

a

It is an average value of the six measurements for each sample.

oligonucleotides C2 , C3 and C4 where each one was at 1.0 × 10−10 M were prepared, and certain amounts of C1 were added. The amounts of C1 in the mixtures were measured using the developed method. The results presented in Table 2 indicated that the developed approach had high accuracy in measuring MDR1 gene level in the MDR1 DNA samples. These observations substantially demonstrate that the method could be potentially used for detection of MDR1 gene in real clinical samples.

4. Conclusions This work has designed a novel electrochemical biosensor platform for rapid determination of MDR1 gene in disease diagnostics by employing Au NPs/TB–GO nanocomposites film modified electrode. The DNA biosensor was label-free, requiring no labeling process or external indicators. Under the optimal conditions, the decline of the peak current was linearly related to the logarithm of the concentration of the target DNA from 1.0 × 10−11 M to 1.0 × 10−9 M, with a detection limit of 2.95 × 10−12 M. The developed method has an ability to discriminate the MDR1 related DNA sequence from even single-base mismatched DNA sequence, to assay the MDR1 related gene level precisely in the MDR1 DNA samples. To fully assess the application potential and added value of the electrochemical DNA biosensor, future research should focus on other different target sequences and sample types, such as the samples derived from cells.

Acknowledgments We gratefully acknowledge the financial support of the National High Technology and Development of China (863 Project: 2012AA022604), the National Natural Science Foundation of China (21275028, 21405015), the Research Fund for the Doctoral Program of Higher Education of China (20123518110001), the Fujian Provincial Important Science and Technology Foundation (2011R1007-2), the Scientific Research Major Program of Fujian Medical University (09ZD013), Sponsored by Medical Elite Cultivation Program of Fujian, P.R.C. (2013-ZQN-JC-5), the Science and Technology Plan Project of General Administration of Quality Supervision (2014IK060), Foundation of Fujian Provincial Department of Education (JA11110, JA12130) and the Natural Science Foundation of Fujian Province of China (2014J05092).

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Biographies Hua-Ping Peng is a lecturer in the Department of Pharmaceutical Analysis in Fujian Medical University. She received her PhD degree in Micro and Nano Materials Science and Engineering from Nanchang University in 2011. Her current research focuses on functional nanomaterials, electrochemical sensor and biosensor. Yan Hu is currently a postgraduate student of Fujian Medical University majoring in pharmaceutical analysis. Pan Liu is currently a postgraduate student of Fujian Medical University majoring in pharmaceutical analysis. Ya-Ni Deng is currently a postgraduate student of Fujian Medical University majoring in pharmaceutical analysis. Peng Wang is currently a postgraduate student of Fujian Medical University majoring in pharmaceutical analysis. Wei Chen is a professor of Pharmaceutical analysis in Fujian Medical University. He received his PhD degree in Analytical Chemistry at Nanjing University in 2010, China. He is also a doctoral supervisor. His research interests are pharmaceutical analysis and nanomaterials in biosensors. Ai-Lin Liu is a professor of Pharmaceutical analysis in Fujian Medical University. He received his PhD degree in Analytical Chemistry from Nanjing University in 2006, China. His current interests include the development of biosensors, chemically modified electrodes and ␮-TAS. Yuan-Zhong Chen is a professor of Fujian Institute of Hematology in the Affiliated Union Hospital of Fujian Medical University. He is a doctoral supervisor. His current interests include the pharmaceutical analysis and chemical biosensors. Xin-Hua Lin is a Professor of Pharmaceutical analysis in Fujian Medical University. He received his MS degrees in Analytical Chemistry from Xiamen University, China, in 1989. He is also a doctoral supervisor. His current research interests cover analytical chemistry, bioelectrochemistry and chemical biosensors.