A reduced graphene oxide-Au based electrochemical biosensor for ultrasensitive detection of enzymatic activity of botulinum neurotoxin A

Sensors and Actuators B 220 (2015) 131–137

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

A reduced graphene oxide-Au based electrochemical biosensor for ultrasensitive detection of enzymatic activity of botulinum neurotoxin A Chun-Yu Chan a , Jiubiao Guo b , Cheng Sun a , Ming-Kiu Tsang c , Feng Tian a , Jianhua Hao c , Sheng Chen b,∗∗ , Mo Yang a,∗ a

Interdisciplinary Division of Biomedical Engineering, the Hong Kong Polytechnic University, Hung Hom, Kowloon, Hong Kong, China Department of Applied Biology and Chemistry, The Hong Kong Polytechnic University, Hung Hom, Kowloon, Hong Kong, China c Department of Applied Physics, the Hong Kong Polytechnic University, Hung Hom, Kowloon, Hong Kong, China b

a r t i c l e

i n f o

Article history: Received 24 February 2015 Received in revised form 2 May 2015 Accepted 22 May 2015 Available online 31 May 2015 Keywords: Reduced graphene oxide (rGO) Electrochemical biosensor Differential pulse voltammetry (DPV) Botulinum neurotoxin type A (BoNT/A)

a b s t r a c t Botulinum neurotoxin serotype A (BoNT/A) is a type of neurotoxin which is able to cause fatal paralytic illness botulism in a low dosage. Therefore, there is a great need to develop an ultrasensitive bioassay to detect its active state for early diagnostics and prevention. This paper presents a reduced graphene oxide (rGO)/Au electrode based electrochemical biosensor for ultrasensitive detection of BoNT serotype A light chain (BoNT-LcA) protease activity. The fabricated rGO/Au electrode provides a robust and biocompatible platform with enhanced electron transfer capability and large area for peptide immobilization. SNAP25-GFP peptide substrate is firstly immobilized on rGO surface via pyrenebutyric acid (PA) linker. The addition of BoNT-LcA could specifically cut SNAP-25-GFP at the cleavage sites to release the cut section from the electrode surface. This enzymatic activity of BoNT-LcA on SNAP-25-GFP peptide substrate could be detected by monitoring the enhanced redox probe transfer rate by differential pulse voltammetry (DPV) with a linear detection range from 1 pg/mL to 1 ng/mL and the limit of detection (LOD) for BoNTLcA is around 8.6 pg/mL. The specificity of this biosensor is demonstrated with BoNT-LcB and heat-treated BoNT-LcA. Moreover, the experiments for BoNT-LcA detection in milk samples demonstrate the feasibility of this biosensor in complex matrix. © 2015 Published by Elsevier B.V.

1. Introduction Bacterial toxins generally are stable proteins secreted by bacterial pathogens, which are the primary virulence factors accounting for infection in permissive hosts [1,2]. Some bacterial toxins such as botulinum neurotoxin (BoNT) are very toxic to humans and considered as potential biological warfare agents [3]. BoNT is a lethal protein toxin which is secreted by Clostridium botulinum [4]. Currently, there are seven identified serotypes (A to G) of botulinum toxins. Among them, BoNT serotype A (BoNT/A) causes the most foodborne botulism events in human [5]. Most of the patients are infected by ingesting improperly handled food. BoNT/A consists of a light chain (LcA) and a heavy chain. The neurotoxicity of BoNT/A

∗ Corresponding author. Tel.: +852 2766 4946; fax: +852 2334 2429. ∗∗ Corresponding author. Tel.: +852 34008795; fax: +852 2364 9932. E-mail addresses: [email protected] (S. Chen), [email protected] (M. Yang). http://dx.doi.org/10.1016/j.snb.2015.05.052 0925-4005/© 2015 Published by Elsevier B.V.

is due to the specific cleavage of (synaptosomal-associated protein 25) SNAP-25 peptide by LcA [6]. Therefore, reliable, rapid and sensitive detection of BoNT-LcA protease activity is important for food safety and botulism prevention. The conventional method for detection of BoNT/A is mousebioassay, which is expensive, labor-intensive, and time-consuming (a minimum 2–3 days for confirmation) with the limit of detection (LOD) around 20 pg/mL for BoNT/A [7]. The alternative methods include enzyme-linked immunosorbent assay (ELISA) [8], fluorescent immunoassay [9] and immunoaffinity column [10,11], which are sensitive (LOD around 10 pg/mL) for laboratory samples, but suffer from complicated experiment protocols and the need of skilled research personnel for operation. Electrochemical biosensors have been proposed for BoNT/A detection, which could provide simple, rapid and low-cost platforms for bacterial toxin detection [12,13]. Limit of detection (LOD) of the current gold electrode based electrochemical biosensor is around tens of pg/mL after peroxidase signal amplification [12,13]. However, the enzymatic signal amplification approach is an indirect detection technique.

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Enzymatic signal amplification used enzyme reactions to amplify the sensing signals. However, enzymes are vulnerable to degradation in the environment to lose their catalytic capabilities. Moreover, the preparation of enzyme is time-consuming and expensive [14]. Some peroxidases such as horseradish peroxidase (HRP), have shown to be carcinogenic to cause serious ecological problems [15,16]. The sensitivity of BoNT/A electrochemical biosensor should also be further improved without peroxidase signal amplification. Therefore, it is necessary to develop a novel electrochemical biosensor for direct detection of BoNT/A with high sensitivity. For an electrochemical biosensor, one of the factors that affect the redox reaction rate in an electrochemical cell is the catalytic ability of the electrode. An electrode with good catalytic ability enhances the redox reaction thus amplifies the signal (current). Graphene has become one of the research focuses in electrochemistry due to its extraordinary electronic property, electrochemical catalytic ability, abundant functional groups for further modification and biocompatibility [17]. Graphene is an excellent electrode candidate in an electrochemical system due to its fast electron transfer rate and electrocatalytic activity [18]. It has been reported that the electron transfer and electrocatalytic activity at the edge of graphene is more efficient than the basal plane. It could be attributed to the high density of electronic states, abundant defective site and oxygen containing groups at the edge of graphene [19]. Various graphene based electrochemical sensors have been developed for detection of mercury ions [20], hydrogen peroxide [21], vascular endothelial growth factor [22], dopamine [23] and DNA hybridization [24]. However, the graphene based electrochemical sensor has not been explored for bacterial protein toxin detection. In this project, a reduced graphene oxide (rGO)/Au based electrochemical biosensor is developed for ultrasensitive detection of enzymatic activity of BoNT-LcA. The rGO/Au electrode surface is firstly functionalized with SNAP-25-GFP substrates via pyrenebutyric acid linker. SNAP-25-GFP layer provides steric hindrance and electrostatic repulsion which decreases the electrochemical signal generated from the working electrode. Detection of BoNT-LcA is achieved by its specific cleavage of SNAP-25-GFP substrate which decreases the steric hindrance and electrostatic repulsion, leading to the recovery of the electrochemical signal. A low limit of detection (LOD) around 5 pg/mL is achieved for BoNT-LcA detection using this rGO/Au electrochemical biosensor without peroxidase signal amplification, which is superior to mouse-bioassay and many immune-based approaches. The specificity of this biosensor was demonstrated using BoNT-LcB and heat-treated LcA. The feasibility of this biosensor for BoNT-LcA detection in milk samples are also explored. Generally, this electrochemical biosensor provides a simple, direct and sensitive platform for detection of enzymatic activity of BoNT/A, which could also be easily adapted to other bacterial toxins detection.

2. Materials and methods 2.1. Materials Graphene oxide synthesized by modified Hummer’s method was purchased from Graphene Supermarket. Potassium hexacyanoferrate (III), potassium hexacyanoferrate(II) trihydrate, hydrazine monohydrate, pyrenebutyric acid (PA), N-hydroxysulfosuccinimide (Sulfo-NHS), MES hydrate, tris(hydroxymethyl)aminomethane, phosphate buffered saline (PBS), sodium dodecyl sulfate (SDS), dl-dithiothreitol (DTT), N-(3-dimethylaminopropyl)-N -ethylcarbodiimide hydrochloride (EDC), and Tween 20 were purchased from Sigma–Aldrich. HEPES (4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid) buffer and

potassium chloride were also obtained from Sigma–Aldrich. PDMS elastomer (Sylgard 184) was purchased from Dow Corning N,N-dimethylformamide (DMF) was purchased from Acros. BoNT LcA and SNAP-25-GFP were constructed according to the previous study [25]. 2.2. Preparation of rGO/Au working electrode A Ti/Au film was firstly deposited on a glass slide by magnetron sputtering. PDMS membrane with a 5 mm diameter circular reservoir was placed on the Au electrode and sealed by silicon rubber in order to define the surface area of working electrode. A metal wire was bonded to the Au electrode (outside the PDMS membrane) using silver paste for connection to an Electrochemical Analyzer (VersaSTAT 3). Finally, the rest of the Au electrode was covered by silicon rubber for passivation. Then, GO solution (0.5 mg/mL, 60 ␮L) was dropped onto the Au electrode and allowed to dry at room temperature. Hydrazine reduction was performed overnight at 65 ◦ C overnight to obtain reduced graphene oxide (rGO). Before the freshly prepared electrode was used, it was activated in 0.3 M KCl by scanning from −0.2 V to 1.05 V with scan rate of 100 mV/s for three CV cycles. 2.3. SNAP-25-peptide conjugation with rGO/Au working electrode To functionalize the rGO surface with SNAP-25-GFP, a heterobifunctional crosslinker–pyrenebutyric acid (PA) was utilized. Firstly, 30 ␮L of 10 mM PA in N,N-dimethylformamide (DMF) solution was incubated with the rGO surface for 2 h. After incubation, the remained PA was washed away with alcohol and DI water sequentially. Finally, 40 ␮L of 10 ␮M SNAP-25-GFP was immobilized on the PA-treated rGO via EDC/NHS chemistry. Excess reactant was washed away by DI water and then 50 mM Tris buffer was added to quench the unreacted NHS ester. The substrate was eventually washed and passivated by 0.5% Tween-20. 2.4. BoNT-LcA enzymatic activity detection and electrochemical analysis BoNT-LcA enzymatic activity detection was performed by adding various concentrations of BoNT LcA toxin solution diluted from a frozen stock solution to the functionalized rGO/Au working electrode. Electrochemical analysis was performed with one conventional three-electrode system with a Pt wire counter electrode, an Ag/AgCl reference electrode and a rGO/Au/SNAP-25-GFP working electrode via a VersaSTAT3 electrochemical analyzer (Princeton Applied Research). Those electrodes were immersed in 1X PBS with 5 mM [Fe(CN)6 ]3−/4− (1:1) and 100 mM KCl for cyclic voltammetry (CV), differential pulse voltammetry (DPV) and electrochemical impedance spectroscopy (EIS) measurement. CV was scanned from −0.2 V to 0.6 V with a scan rate of 50 mV/s. DPV was carried out with scanning range from 0.1 V to 0.5 V, 0.02 s pulse width, 4 mV/s scan rate and 50 mV pulse height. EIS was performed with scan frequency ranged from 0.1 Hz to 100 kHz, potential of 0.2 V and amplitude of 5 mV. To quantitatively determinate LcA concentration, sensor response was represented by relative change in peak current of DPV (I%). 2.5. Characterization TEM images were taken by a JEOL-2100F transmission electron microscopy (TEM) equipped with an Oxford Instrument EDS system, operating at 200 kV. Raman spectra of GO and rGO was recorded by a Horiba HR800 Raman spectrometer. XPS spectra were

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Fig. 1. Schematic illustration of the detection mechanism of the rGO based biosensors.

taken with a SKL-12 spectrometer modified with VG CLAM 4 multichannel hemispherical analyzer. 3. Results

complex, conjugation of SNAP-25-GFP peptide on rGO was confirmed by the appearance of characteristic peaks corresponding to the CH2 stretching vibration (2925 cm−1 and 2860 cm−1 ), amide group vibration (1660 cm−1 and 1581 cm−1 ) and C O stretching vibration (1213 cm−1 ).

3.1. Sensing mechanism The sensing mechanism of the proposed rGO/Au electrochemical biosensor for detecting enzymatic activity of botulinum neurotoxin A is shown in Fig. 1. Initially, pyrenebutyric acid with aromatic rings is conjugated with rGO surface by ␲–␲ stacking. SNAP-25-GFP peptide is then immobilized on rGO surface via pyrenebutyric acid linker by EDC/NHS activation. The added target BoNT-LcAs specifically cut immobilized SNAP-25-GFP at the cleavage site and the cut section then detach from rGO/Au electrode surface. Due to the decreasing hindrance of redox probes transfer on electrodes, electrochemical currents increased which can be used for BoNT-LcA enzymatic activity detection. 3.2. Characterization of rGO and peptide conjugation After reduction, rGO was generated from GO sheets. Fig. 2a shows the TEM image of rGO sheets on substrates with the average size of 500 nm. Ripples and wrinkles were observed on rGO surface in high resolution TEM image which could largely increase the contact area with redox probes (Fig. 2b). Reduction process of GO was investigated by Raman Spectroscopy (Fig. 2c). Graphene oxide has two signature peaks in Raman spectroscopy: (a) D band at ∼1350 cm−1 and (b) G band at ∼1600 cm−1 . Before reduction, relative intensity of G band in GO was higher than that of D band. After hydrazine reduction, relative intensity of D band in rGO exceeded the one of G band. The conjugation of SNAP-25-GFP on rGO was then characterized by both XPS and FTIR. As shown in Fig. 3a, there was an obvious nitrogen peak (N1s ) appearing in XPS spectrum of rGO/SNAP25-GFP which was attributed to the nitrogen components from SNAP-25-GFP grafted on rGO surface. Fig. 3b shows the FTIR spectra of rGO before and after conjugation with SNAP-25-GFP. Before conjugation, a small characteristic peak of rGO for O H was observed at 3401 cm−1 . Other oxidative peaks including C O and C O could not be observed due to the reduction process. For the rGO–peptide

3.3. Establishment of rGO/Au electrode with grafted peptide To demonstrate the formation of rGO/Au electrode and the assembly of SNAP-25-GFP peptide, cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS) were used to characterize the assembly steps. A reduction peak and a oxidation peak of [Fe(CN)6 ]3−/4− redox probes can be obviously observed for the bare Au electrode (Fig. 4, curve a). After the assembly of rGO on Au electrode surface, there is an obvious increase of peak current for both redox peaks (Fig. 4, curve b). The peak current obviously decreased after PA grafting on rGO surface via ␲–␲ stacking (Fig. 4, curve c). It could be due to the electrostatic repulsion between redox probe [Fe(CN)6 ]3−/4− and the deprotonated carboxyl group ( COO− ) in PA. When SNPA-25-GFP was conjugated to PA through EDC/NHS approach, the peak current was further decreased (Fig. 4, curve d). The peak current increase from curve a (Au) to curve b (Au/rGO) could be attributed to the superior catalytic ability and charge transfer ability of graphene. To demonstrate this, cathodic and anodic peak currents (ipa and ipc ) were measured from the rGO/Au electrode. As shown in Fig. 5, ipa and ipc increased with the scan rate and a linear relationship between ipa or ipc and square root of scan rate was observed, which demonstrated that the electrochemical reaction of electrodes was due to diffusion of redox probes with a solution phase quasi-reversible process. For the peak current decrease from curve c to curve d, there were two major factors contributed to this observation. Firstly, coverage of SNAP-25-GFP on the rGO surface increased steric hindrance which prevented the redox probe performing electron-transfer on the electrode surface. Secondly, the isoelectric point of SNAP-25-GFP was lower than the pH value of the buffer solution (pH 7.4). SNAP-25-GFP was deprotonated with negative charge which provided electrostatic repulsion force which further decreased the electrochemical signal by repelling the redox probe away from electrode surface.

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(a) C1s

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Wave number (cm-1) Fig. 3. (a) XPS spectra of rGO and rGO/PA/SNAP-25-GFP; (b) FTIR spectra of rGO and rGO/PA/SNAP-25-GFP.

EIS was also used to measure the impedance of electrodes over a range of frequencies during surface modification steps. The data of EIS was expressed as Nyquist plot (Fig. 6). The semicircle of Nyquist plot at high frequency range represented the electron-transfer process at the electrode surface. Charge transfer resistance (Rct ) of the electrode can be obtained from the diameter of the semicircle. The change of Rct in EIS spectrum generally agreed with our observation in CV scanning. The coating of rGO on Au electrode enhanced the electron transfer process which reduced the Rct . Functionalization of PA and SNAP-25 shielded the electrode surface from redox probe by steric hindrance and electrostatic repulsion, which increased the Rct .

300 b 200

Current (μA)

Fig. 2. (a) TEM image of rGO flakes; (b) high resolution TEM image of rGO sheets with ripples and wrinkles; (c) Raman Spectroscopy of GO before and after reduction.

biosensor for LcA detection, DPV was chosen as the sensing signal instead of CV measurement. The reason is that CV measurement shows hysteresis effect due to the double layer capacitance at the electrode surface leading to a capacitive current [28]. In contrast, pulse techniques such as differential pulse voltammetry (DPV) provide the highest sensitivities among electrochemical methods with the ability to subtract the double-layer charging current. The double layer charging current will be filtered by pulse techniques to avoid the hysteresis effect and provide quantitative study in low concentrations [29,30]. Therefore, DPV was chosen as the sensing signal in our electrochemical sensor for LcA detection. Fig. 7a shows

c a

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3.4. LcA detection by differential pulse voltammetry (DPV)

-100

LcA is the proteolytic active subunit in BoNT/A and was reported to be a Zn2+ dependent metalloprotease [26]. Zn2+ (10 ␮M) and DTT (2 mM) in the buffer solution is crucial for maintaining LcA activity. In many instances of electrical biosensing devices such as field effect transistor (FET), hysteresis effect will cause a lag in the response during electrical scans which may screen the sensing signals [27]. To evaluate the performance of this electrochemical

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Potential (V) Fig. 4. Cyclic voltammograms of Au, rGO/Au, Au/rGO/PA and Au/rGO/PA/SNAP-25GFP modified electrodes.

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(a) 100 Blank buffer

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Potential (V vs Ag/AgCl) Fig. 5. Cyclic voltammogram of Au/rGO electrode scanning from −0.2 V to 0.6 V with various scan rate of (10 mV/s, 20 mV/s, 40 mV/s, 60 mV/s, 80 mV/s, 100 mV/s) in pH 7.4 PBS supplemented with 5 mM [Fe(CN)6 ]3−/4− and 100 mM KCl. Inset plots the anodic (ipa ) and cathodic current (ipc ) against the square root of scanning rate.

(b)

70 60

y = 9.0945ln(x) - 6.4527 R² = 0.9802

typical DPV response curves of Au/rGO/PA/SNAP-25-GFP electrode after incubating with 1 ng/mL LcA for 1 h. A well-defined DPV peak was recorded with the scanning from 0.1 V to 0.5 V. As shown, the peak current increased with addition of LcA after incubation which was due to the specific cleavage of SNAP-25-GFP peptide immobilized on electrode surface. The enzymatic activity of BoNT-LcA was reflected by the magnitude change of the peak currents. The peak current increase could be explained by the proteolytic activity on SNAP-25-GFP substrates. LcA cleaved SNAP-25-GFP immobilized on electrodes during incubation, which decreased the electrostatic repulsion and steric hindrance effect on the electrode. When LcA was introduced, the proteolytic activity of LcA cleaved the GFP section of SNAP-25-GFP from the rGO electrode. It reduced the size of peptide immobilized on rGO electrode and decreased the coverage degree of electrode surface, which made the redox charge transfer to electrode surface much easier, i.e. steric hindrance was decreased. Moreover, the calculated pKa of cleaved part of GFP section was less than 7 which meant the removed GFP section was more negative charged compared with the remained section of SNAP-25-GFP in neutral buffer. This cleavage decreased the charge density on rGO surface and made the electrode surface less negative, which would let the negatively charged redox probes much easier approach the electrode surface to increase the electrochemical signals. Thus, the electron transfer enhanced which led to the peak current increase of DPV curves. Fig. 7b shows the relative peak current change (I%) with various concentrations of LcA. It was observed that the peak current change (I%) of this biosensor

ΔI %

50 40 30 20 10 0 1

10

100

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LcA concentration (pg/ml) Fig. 7. (a) DPV curves of functionalized rGO/Au electrodes with control buffer and fresh BoNT-LcA of 1 ng/mL; (b) calibration curve for the change in DPV current (I%) at different BoNT-LcA concentrations.

is linearly increased with the logarithm concentrations of BoNTLcA in the range from 1 pg/mL to 1 ng/mL. The linear equation is I% = −10.168 + 22.848 × log (LcA) with R2 = 0.997. This logarithm correlation curve between sensing signals and toxin concentrations match the results from literatures [31–33]. This could be explained by the catalytic nature of LcA in proteolytic activity. Generally, LcA molecule could continuously cleave many SNAP25-GFP peptides without losing its proteolytic capability. In ideal case, one LcA molecule could even cleave all the SNAP-25-GFP peptides. Therefore, the increase of LcA concentrations will not lead to a corresponding linear increase of sensing signals. Instead, the peak current change of the biosensor was linearly increased with the logarithm concentration of BoNT-LcA. 3.5. Specificity testing

50

-Z" (Ω)

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Au/rGO/PA Au/rGO/PA/peptide

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Z' (Ω) Fig. 6. Nyquist plot with Au, rGO/Au, Au/rGO/PA and Au/rGO/PA/SNAP-25-GFP modified electrodes.

To testify the specificity of this biosensor for BoNT-LcA detection, BoNT serotype B light chain (BoNT-LcB) (1 ng/mL) and heat-treated LcA (1 ng/mL) were tested with identical procedures with fresh LcA. BoNT-LcB was also toxic to humans but could not cleave SNAP-25 substrate and heat-treated LcA was denatured to lose enzyme capability after high temperature treatment. As shown in Fig. 8a, there were no obvious difference among the peak currents of DPV curves for control buffer, LcB and heat-treated LcA because both LcB and heat-treated LcA could not cleave SNAP-25GFP substrates. In contrast, fresh LcA with the same concentration significantly increased the peak current due to the specific cleavage on SNAP-25-GFP substrates. Fig. 8b shows the relative peak current change (I%) for control sample, LcB, heat-treated LcA and fresh LcA with the same concentration of 1 ng/mL. It was obviously that the relative peak current change (I%) for LcA was much larger

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Fig. 9. Comparison of the rGO/Au electrochemical biosensors for BoNT-LcA detection with various concentrations in buffer and skimmed milk samples.

80

any non-specifically assorted proteins. In such situation, SNAP-25GFP was not cleaved and non-specific adsorption was eliminated, milk was found to have a small interference on our sensor’s performance. The limit of detection (LOD) was calculated as 8.6 pg/mL based on the background signal plus 3 times of standard derivation. The above results demonstrated the potential applicability of this biosensor for real food sample detection.

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ΔI %

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4. Conclusion

10 0 Control Noise

LcB

Heat-treated LcA

LcA

Fig. 8. (a) Specificity testing for DPV curves of functionalized rGO/Au electrodes with control buffer, BoNT-LcB (1 ng/mL), heat-treated BoNT-LcA (1 ng/mL) and fresh BoNT-LcA (1 ng/mL); (b) statistical analysis of relative DPV peak current change (I%) for various samples incubation.

than LcB and heat-treated LcA, which demonstrated the specificity of this biosensor for LcA detection. 3.6. LcA detection in milk When biosensors work in real samples, the variance and fluctuation of signals caused by interference of various components in complex matrix may screen the sensing signals [34,35]. It is necessary to study the performance of biosensors in complex matrix to study the interference effect. Since foodborne is the major infection rout of BoNT-LcA, it is crucial to evaluate the performance of this biosensor in simulated food samples. BoNT/A is a protein toxin whose activity is mainly interfered by other protein components in real samples due to non-specific reaction. Since milk is full of proteins, it is always used as complex protein abundant matrix to study protein components interference for BoNT detection [36]. For this purpose, BoNT-LcA protease activity detection was explored in skimmed milk spiked with various concentrations of BoNT-LcA. The other conditions were the same with previous experiments in buffer solution. As shown in Fig. 9, the relative peak current changes (I%) for toxin spiked skimmed milk for all the concentrations were slightly decreased compared with those in buffer, which demonstrated that the interference from abundant proteins in the complex environment of skimmed milk had a very small effect on the sensing signals. Although there were abundant amount of proteins in milk, those proteins do not have any proteolytic activity against our detection probe (SNAP-25-GFP). The only possibility that the sensor would be affected was non-specific adsorption. This possible non-specific adsorption was avoided by a washing step with Tween-20 after sample incubation to effectively remove

In this paper, a reduced graphene oxide/Au electrochemical biosensor was developed for detection of enzymatic activity of botulinum neurotoxin (BoNT-LcA). The functionalization of SNAP25-GFP substrates on rGO/Au electrode surface was realized by pyrenebutyric acid linker. The addition of BoNT-LcA led to the cleavage of SNAP-25-GFP substrates at the cleavage sites which could be monitored by electrochemical signal change. The feasibility of this rGO/Au electrochemical biosensor for BoNT-LcA detection was demonstrated with a linear working range from 1 pg/mL to 1 ng/mL and the LOD of 8.6 pg/mL. The experiments with BoNT-LcB and heat-treated BoNT-LcA demonstrated the good specificity of this biosensor. The experiments in skimmed milk also demonstrated the feasibility of this biosensor in complex matrix. Acknowledgement This work was supported by the internal research funds of the Hong Kong Polytechnic University (G-YN03 and G-YBAA). References [1] D.M. Gill, Bacterial toxins: a table of lethal amounts, Microbiol. Rev. 46 (1982) 86–94. [2] R. Rappuoli, M. Pizza, G. Douce, G. Dougan, New vaccines against bacterial toxins, Adv. Exp. Med. Biol. 397 (1996) 55–60. [3] G.W. Christopher, T.J. Cieslak, J.A. Pavlin, E.M. Eitzen, Biological warfare. A historical perspective, J. Am. Med. Assoc. 278 (1997) 412–417. [4] B.R. DasGupta, J. Foley Jr., C. botulinum neurotoxin types A and E: isolated light chain breaks down into two fragments. Comparison of their amino acid sequences with tetanus neurotoxin, Biochimie 71 (1989) 1193–1200. [5] J. Sobel, N. Tucker, A. Sulka, J. McLaughlin, S. Maslanka, Foodborne botulism in the United States, 1990–2000, Emerg. Infect. Dis. 10 (2004) 1606–1611. [6] P.G. Foran, N. Mohammed, G.O. Lisk, S. Nagwaney, G.W. Lawrence, E. Johnson, L. Smith, K.R. Aoki, J.O. Dolly, Evaluation of the therapeutic usefulness of botulinum neurotoxin B, C1, E, and F compared with the long lasting type A. Basis for distinct durations of inhibition of exocytosis in central neurons, J. Biol. Chem. 278 (2003) 1363–1371. [7] R.L. Shapiro, C. Hatheway, D.L. Swerdlow, Botulism in the United States: a clinical and epidemiologic review, Ann. Intern. Med. 129 (1998) 221–228. [8] S. Bok, V. Korampally, C.M. Darr, W.R. Folk, L. Polo-Parada, K. Gangopadhyay, S. Gangopadhyay, Femtogram-level detection of Clostridium botulinum neurotoxin type A by sandwich immunoassay using nanoporous substrate and ultra-bright fluorescent suprananoparticles, Biosens. Bioelectron. 41 (2013) 409–416.

C.-Y. Chan et al. / Sensors and Actuators B 220 (2015) 131–137 [9] R.G. Jones, M. Ochiai, Y. Liu, T. Ekong, D. Sesardic, Development of improved SNAP25 endopeptidase immuno-assays for botulinum type A and E toxins, J. Immunol. Methods 329 (2008) 92–101. [10] F. Gessler, S. Pagel-Wieder, M.A. Avondet, H. Böhnel, A fluorescence detection platform using spatial electroluminescent excitation for measuring botulinum neurotoxin A activity, Diagn. Microbiol. Infect. Dis. 57 (2007) 243–249. [11] O. Attrée, V. Guglielmo-Viret, V. Gros, P. Thullier, Development and comparison of two immunoassay formats for rapid detection of botulinum neurotoxin type A, J. Immunol. Methods 325 (2007) 78–87. [12] F. Wei, C.M. Ho, Aptamer-based electrochemical biosensor for botulinum neurotoxin, Anal. Bioanal. Chem. 393 (2009) 1943–1948. [13] F. Wei, B. Bai, C.M. Ho, Rapidly optimizing an aptamer based BoNT sensor by feedback system control (FSC) scheme, Biosens. Bioelectron. 15 (2011) 174–179. [14] T. Zheng, Q. Zhang, S. Feng, J.J. Zhu, Q. Wang, H. Wang, Robust nonenzymatic hybrid nanoelectrocatalysts for signal amplification toward ultrasensitive electrochemical cytosensing, J. Am. Chem. Soc. 136 (2014) 2288–2291. [15] R.A. Floyd, L.M. Soong, P.L. Culver, Horseradish peroxidase/hydrogen peroxidecatalyzed oxidation of the carcinogen N-hydroxy-N-acetyl-2-aminofluorene as effected by cyanide and ascorbate, Cancer Res. 36 (1976) 1510–1519. [16] M. Stiborová, B. Asfaw, P. Anzenbacher, Activation of carcinogens by peroxidase horseradish peroxidase-mediated formation of benzenediazonium ion from a non-aminoazo dye, 1-phenylazo-2-hydroxynaphthalene (Sudan I) and its binding to DNA, FEBS Lett. 232 (1988) 387–390. [17] X.W. Mao, G.C. Rutledge, T.A. Hatton, Nanocarbon-based electrochemical systems for sensing, electrocatalysis, and energy storage, Nano Today 9 (2014) 405–432. [18] R.L. McCreery, Advanced carbon electrode materials for molecular electrochemistry, Chem. Rev. 108 (2008) 2646–2687. [19] W. Yuan, Y. Zhou, Y. Li, C. Li, H. Peng, J. Zhang, Z. Liu, L. Dai, G. Shi, The edgeand basal-plane-specific electrochemistry of a single-layer graphene sheet, Sci. Rep. 3 (2013) 2248. [20] Y. Zhang, H. Zhao, Z.J. Wu, Y. Xue, X.F. Zhang, Y.J. He, X.J. Li, Z.B. Yuan, A novel graphene-DNA biosensor for selective detection of mercury ions, Biosens. Bioelectron. 48 (2013) 180–187. [21] A.M. Farah, F.T. Thema, E.D. Dikio, Electrochemical detection of hydrogen peroxide based on graphene oxide/prussian blue modified glassy carbon electrode, Int. J. Electrochem. Sci. 7 (2012) 5069–5083. [22] C.W. Lin, K.C. Wei, S.S. Liao, C.Y. Huang, C.L. Sunc, P.J. Wua, Y.J. Lu, H.W. Yang, C.C. Ma, A reusable magnetic graphene oxide-modified biosensor for vascular endothelial growth factor detection in cancer diagnosis, Biosens. Bioelectron. 15 (2015) 431–437. [23] A.T.E. Vilian, S.M. Chen, Simple approach for the immobilization of horseradish peroxidase on poly-l-histidine modified reduced graphene oxide for amperometric determination of dopamine and H2 O2 , RSC Adv. 4 (2014) 55867–55876. [24] A. Bonanni, M. Pumera, Graphene platform for hairpin-DNA-based impedimetric genosensing, ACS Nano 5 (2011) 2356–2361. [25] W.W. Ye, J.B. Guo, S. Chen, M. Yang, Nanoporous membrane based impedance sensors to detect the enzymatic activity of botulinum neurotoxin A, J. Mater. Chem. B 1 (2013) 6544–6550. [26] G. Schiavo, O. Rossetto, A. Santucci, B.R. DasGupta, C. Montecucco, Botulinum neurotoxins serotypes A and E cleave SNAP-25 at distinct COOH-terminal peptide bonds, J. Biol. Chem. 267 (1992) 23479–23483. [27] Y. Paska, H. Haick, Interactive effect of hysteresis and surface chemistry on gated silicon nanowire gas sensors, ACS Appl. Mater. Interfaces 4 (2012) 2604–2617. [28] R.O. Ansell, A. Mcnaughtan, Electrochemical sensors, in: M. Campbell (Ed.), Sensor Systems for Environmental Monitoring, Blackie Academic & Professional Inc., London, 1997, pp. 100–126. [29] J. Koehne, J. Li, A.M. Cassell, H. Chen, Q. Ye, H.T. Ng, J. Han, M. Meyyappan, J. Mater. Chem. 14 (2004) 676–684. [30] F. Zhao, R.C.T. Slade, J.R. Varcoe, Techniques for the study and development of microbial fuel cells: an electrochemical perspective, Chem. Soc. Rev. 38 (2009) 1926–1939.

137

[31] K. Bagramyan, J.R. Barash, S.S. Arnon, M. Kalkum, Attomolar detection of botulinum toxin type A in complex biological matrices, PLoS One 3 (2008) e2041. [32] M.L. Frisk, E. Berthier, W.H. Tepp, E.A. Johnson, D.J. Beebe, Bead-based microfluidic toxin sensor integrating evaporative signal amplification, Lab Chip 8 (2008) 1793–1800. [33] G.B. Stevens, D.A. Silver, A. Zgaga-Griesz, W.G. Bessler, S.K. Vashist, P. Patel, K. Achazi, J. Strotmeier, S. Worbs, M.B. Dorner, B.G. Dorner, D. Pauly, A. Rummel, G.A. Urban, M. Krueger, Bioluminescence assay for the highly sensitive detection of botulinum neurotoxin A activity, Analyst 138 (2013) 6154–6162. [34] N. Shehada, G. Bronstrup, K. Funka, S. Christiansen, M. Leja, H. Haick, Ultrasensitive silicon nanowire for real-world gas sensing: noninvasive diagnosis of cancer from breath volatolome, Chem. Rev. 112 (2012) 5949–5966. [35] B. Wang, J.C. Cancilla, J.S. Torrecilla, H. Haick, B. Wang, Artificial sensing intelligence with silicon nanowires for ultraselective detection in the gas phase, Nano Lett. 14 (14) (2014) 933–938. [36] O.G. Weingart, T. Schreiber, C. Mascher, D. Pauly, M.B. Dorner, T.F.H. Berger, C. Egger, F. Gessler, M.J. Loessner, M.A. Avondet, Appl. Environ. Microbiol. 76 (76) (2010) 3293–3300.

Biographies Chun-Yu Chan received B.S. degree from Hong Kong University of Science and Technology in 2009 and M.Sc. degree from the Hong Kong Polytechnic University in 2012. Currently, he is working toward the Ph.D. degree in the Interdisciplinary Division of Biomedical Engineering of the Hong Kong Polytechnic University since 2012. His research interests include nanomaterials based electrochemical biosensors. Jiubiao Guo received M.S. degree from Tsinghua University, in 2010. He is currently a Ph.D. student in Department of Applied Biology and Chemical Technology in the Hong Kong Polytechnic University since 2012. His research interests include bacterial toxin detection. Cheng Sun received B.S. degree from the North University of China in 2013. He is currently a M.Sc. student in the Interdisciplinary Division of Biomedical Engineering of the Hong Kong Polytechnic University. His research interests include graphene based electrochemical biosensor. Ming-Kiu Tsang obtained his B.Sc. degree in 2011 from the Hong Kong Polytechnic University. Then he received his M.Phil. degree in 2013 from the Hong Kong Polytechnic University. After graduation, he continued to pursue his Ph.D. in Department of Applied Physics. His current research interests include lanthanide doped nanomaterials for bio-medical imaging, bio-detection and therapeutic applications. Jianhua Hao obtained his B.Sc., M.Sc. and Ph.D. from Huazhong University of Science and Technology, China. He is now an associate professor in the Department of Applied Physics in the Hong Kong Polytechnic University. His research interests include luminescent materials for photonic and biological applications, thin-films and heterostructures and nanomaterials. Sheng Chen obtained his B.S. and M.S. degree from China Agriculture University in 1997 and 2000, respectively. He then obtained his Ph.D. degree from University of Maryland at College Park in 2004. He is currently an associate professor in Department of Applied Biology and Chemical Technology in the Hong Kong Polytechnic University. His research interests include molecular mechanisms of antimicrobial resistance and bacterial toxin detection. Mo Yang received the B.E. and M.S. degrees in power mechanical engineering from Shanghai Jiaotong University, Shanghai, China, in 1998 and 2001, respectively. He received the Ph.D. degree in mechanical engineering from the University of California, Riverside, in 2004. Currently, he is an associate professor in Interdisciplinary Division of Biomedical Engineering of the Hong Kong Polytechnic University. His research interests include nanomaterials based biosensor.