Biosensors and Bioelectronics 70 (2015) 418–425
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Ultrasensitive detection of drug resistant cancer cells in biological matrixes using an amperometric nanobiosensor Pranjal Chandra a,b, Hui-Bog Noh a, Ramjee Pallela a, Yoon-Bo Shim a,n a b
Department of Chemistry and Institute of Biophysio Sensor Technology (IBST), Pusan National University, Busan 609-735, South Korea Amity Institute of Biotechnology, Amity University Uttar Pradesh, Noida 201303, U.P., India
art ic l e i nf o
a b s t r a c t
Article history: Received 5 February 2015 Received in revised form 10 March 2015 Accepted 26 March 2015 Available online 27 March 2015
Multidrug resistance (MDR) is a key issue in the failure of cancer chemotherapy and its detection will be helpful to develop suitable therapeutic strategies for cancer patients and overcome the death rates. In this direction, we designed a new amperometric sensor (a medical device prototype) to detect drug resistant cancer cells by sensing “Permeability glycoprotein (P-gp)”. The sensor probe is fabricated by immobilizing monoclonal P-gp antibody on the gold nanoparticles (AuNPs) conducting polymer composite. The detection relies on a sandwich-type approach using a bioconjugate, where the aminophenyl boronic acid (APBA) served as a recognition molecule which binds with the cell surface glycans and hydrazine (Hyd) served as an electrocatalyst for the reduction of H2O2 which are attached on multi-wall carbon nanotube (MWCNT) (APBA-MWCNT-Hyd). A linear range for the cancer cell detection is obtained between 50 and 100,000 cells/mL with the detection limit of 237 2 cells/mL. The proposed immunosensor is successfully applied to detect MDR cancer cells (MDRCC) in serum and mixed cell samples. Interferences by drug sensitive (SKBr-3 and HeLa), noncancerous cells (HEK-293 and OSE), and other chemical molecules present in the real sample matrix are examined. The sensitivity of the proposed immunosensor is excellent compared with the conventional reporter antibody based assay. & 2015 Elsevier B.V. All rights reserved.
Keywords: Nanobiosensor Drug resistant cancer cells Electrochemistry In vitro test Nanomaterials
1. Introduction Multidrug resistance (MDR) is a major factor in the failure of cancer chemotherapy. Several mechanisms have been convinced to play vital roles in the development of the MDR in cancer cells (Krishna and Mayer, 2000). Among them, a major form is mediated by a cell membrane transporter “Permeability glycoprotein, (P-gp)” which is encoded by the MDR1 gene in human cells. It includes two nucleotide-binding and two membrane-spanning domains, acts as energy dependent pump that decreases intracellular drug concentrations lower than the effective therapeutic levels (Bellamy, 1996; Gottesman et al., 2002). This phenomenon leads to the development of drug resistance in cancer cells. The early detection of MDR in cancer cells can lead to the adequate chemotherapy of cancer patients and overcome the mortality. Thus, it is extremely important and clinically significant to detect MDRCC at the early stage. This will help the clinicians to develop alternate therapeutic strategies for cancer patients. Until now, several biological methods have been developed to detect P-gp expression for the diagnosis of MDR in cancer such as; n
Corresponding author. Fax: þ82 51 514 2122. E-mail address: [email protected] (Y.-B. Shim).
http://dx.doi.org/10.1016/j.bios.2015.03.069 0956-5663/& 2015 Elsevier B.V. All rights reserved.
polymerase chain reaction (Murphy et al., 1990), immunohistochemistry (Chan et al., 1990), flow cytometry (Ludescher et al., 1992), and microarray (Gillet et al., 2004). Although these methods can be used to detect MDRCC, they are less sensitive, require highly trained professionals, and have no ability to be miniaturized for the point-of-care applications. Thus, few biosensor-based diagnostic methods have also been reported to detect MDRCC. These methods, however, are indirect, nonselective, and less sensitive (Du et al., 2005; Zhang et al., 2011, 2014). Thus, the development of a sensitive, robust, and alternate method for the detection of MDRCC directly in the biological fluids is desirable and has great clinical importance. To achieve a highly sensitive detection method for MDRCC, a sandwich type-electrochemical immunosensor composed of P-gp monoclonal antibody (AntiP-gp) as a detection probe coupled with a ligand tagged-non-enzymatic catalyst as a reporter probe is worthy to attempt due to their selective and highly sensitive characteristics, respectively. The stable immobilization of biomaterials including antibodies on the sensor probe is extremely critical in the biosensor fabrication (Rahman et al., 2008). Electrochemical biosensors composed of conducting polymers-AuNPs composite are considered to be highly stable and ultrasensitive because biomolecules can be covalently immobilized on the polymer backbone possessing –COOH or –NH2 groups (Chandra et al., 2011a, 2013; Koh et al.,
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2011; Kim and Shim, 2013; Lee et al., 2010). Compare to the reporter antibody based conventional sandwiched type electrochemical immunosensor (Malhotra et al. 2012; Wu et al., 2015; Zhu et al. 2014), other ligands are worthy to attempt for the in vitro detection of cancer cells. This is important because it has been reported that the concentration of antigen (target molecule) on the cancer cell surface is low (Beck et al., 1996). This may result in low antigen–antibody binding and consequently poor detection of cancer cells. Thus, it is worthy to target molecules that are present on the cancer cell surface in abundance. Recent studies have clearly indicated existence of high concentration of glycans onto the cancer cell membrane (Dube and Bertozzi, 2005; Zhang et al., 2010) due to which numerous diol-groups are present on the cell surface. Arising from the unique capacity of boronic acids to form boronic esters with these diols (Das et al., 2011), it would be interesting to attempt boronic acid as an alternative recognition molecule (reporter probe) to detect cancer cell surface glycans. To show the signal in an electrochemical immunosensor, an electrochemical indicator is needed, such as a bioconjugate composed of enzyme or nonenzymatic catalyst. Compared to expensive and easily denaturable enzyme, Hyd attached on a bioconjugate can be used due to its small size, stability, and high catalytic activity towards hydrogen peroxide (H2O2) reduction (Zhu et al., 2013). Thus, in this work we developed a sandwich immunoassay approach having a chemo-nano-conjugate composed of APBA, MWCNT, and Hyd instead of enzyme linked reporter antibody. In this conjugate, APBA and Hyd are used a reporter molecule and electrocatalyst for signal generation, respectively. In the present study, a novel amperometric immunosensor for the detection of MDRCC has been tried to develop for the first time through the detection of P-gp. The immunosensor probe is fabricated by forming covalent bonds between AntiP-gp and carboxylic acid group-functionalized conducting polymer layer on AuNPs deposited electrode surface. The nanocomposite was characterized by atomic force microscopy (AFM), X-ray photoelectron spectroscopy (XPS), electrochemical impedance spectroscopy (EIS), scanning electron microscopy (SEM), and the AntiP-gp immobilization was confirmed by quartz crystal microbalance (QCM) and XPS. The sandwich immunosensing approach was adopted where APBAMWCNT-Hyd was reacted with the MDRCC captured by the GCE/ AuNPs/pTTBA/AntiP-gp probe. The immunoreaction was monitored in term of catalytic activity of Hyd towards H2O2 reduction. The experimental parameters were optimized and the detection limit of the MDRCC was determined. Direct detection of MDRCC in serum and in the mixed cell samples were performed to evaluate the real clinical value of sensor. The selectivity of the biosensor was also examined toward various non target cells, and chemicals present in the real sample matrix. The developed method was also compared with the conventional reporter antibody based method.
2. Material and methods 2.1. Materials 2,2′:5′, 2″-terthiophene-3′(p-benzoic acid) (TTBA) was synthesized through the Paal–Knorr pyrrole condensation reaction (Koh et al., 2011). Tetrabutylammonium perchlorate (TBAP, electrochemical grade) was purchased from Fluka (USA) and purified according to a general method, followed by drying under vacuum at 1.33 10 3 Pa (Noh et al., 2012). 1-ethyl-3-(3-(dimethylamino)propyl) carbodiimide (EDC), N-Hydroxysuccinimide (NHS), dichloromethane (99.8%, anhydrous), trisodium citrate, sodium tetrahydridoborate, HAuCl4 3H2O, bovine serum albumin (BSA), and Hydrazine sulfate were purchased from sigma Aldrich (USA).
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Monoclonal p-GP antibody, amino-phenylboronic acid, indium tin oxide (ITO) glass, and H2O2 (33%) were purchased from sigma Aldrich (USA). MWCNTs (4–6 nm diameter, 95%) were obtained from Iljin Nanotech (South Korea). phosphate buffer saline solutions (PBS) were prepared with 0.1 M of disodium Hydrogen phosphate, 0.1 M of sodium dihydrogenphosphate, and 0.9% sodium chloride. cell culture medium, fetal bovine serum (FBS), trypsin-EDTA, penicillin/streptomycin, hank’s balance salt (HBS) solution, were purchased from Sigma-Aldrich (USA). All other chemicals were of extra pure analytical grade and used without further purification. all aqueous solutions were prepared in ultrapure water, which was obtained from a Milli-Q water purifying system (18 M Ω cm). 2.2. Apparatus All electrochemical experiments were performed in a three electrode cell. The modified glassy carbon electrode (GCE) (dia. 3.0 mm), Ag/AgCl (in saturated KCl), and a platinum (Pt) wire were used as working, reference, and counter electrodes, respectively. Voltammograms and chronoamperograms were recorded using a potentiostat/galvanostat, Kosentech, model KST P-2 (South Korea). A Multimode AFM device (Veeco Metrology) equipped with a Nanoscope IV controller (Veeco) was used at ambient conditions to get the images. The QCM experiment was performed using a SEIKO EG&G model QCA 917 and a PAR model 263A potentiostat/ galvanostat (USA). An Au-coated working electrode (area: 0.196 cm2; 9 MHz; AT-cut quartz crystal) was used for the QCM experiment. The impedance spectra were obtained using a EG&G Princeton Applied Research PARSTAT2263 at an open circuit voltage from 100.0 kHz down to 100.0 mHz and a sampling rate of five points per decade (AC amplitude: 10.0 mV). XPS was performed using a VG Scientific XPSLAB 250 XPS spectrometer and a monochromated Al Kα source with charge compensation at KBSI (Busan). The SEM images were obtained using a Cambridge Stereoscan 240. A JEOL JEM-2010 electron microscope (Jeol High-Tech Co., Japan) with an acceleration voltage of 200 kV was used to obtain TEM images. 2.3. Preparation of APBA-MWCNT-Hyd conjugate Firstly, MWCNTs were functionalized according to previously reported methods (Goldman and Lellouche, 2010; Piran et al., 2009). Briefly, 100 mg of MWCNT was treated with mixture of concentrated 12.0 M HNO3 and 36.0 M H2SO4 (90 °C, 2.0 h) followed by multiple rinsing with deionized water until no acid was detected. After drying at 80 °C under vacuum overnight, a black powder was obtained. Next, the treated MWCNT (3.0 mg) was dispersed into 1.0 mL PBS (pH 7.0) containing 10.0 mM EDC/NHS solutions and incubated at room temperature for 6 h to activate the –COOH on the MWCNTs. The resulting mixture was separated by centrifugation and the precipitate was washed three times. Meanwhile, an optimized 1.0 mg/mL hydrazine sulfate solution and 10.0 mM APBA was prepared in PBS. Subsequently, the activated MWCNTs were mixed with hydrazine sulfate and APBA solution. The mixture was stored in a refrigerator overnight at 4 °C followed by centrifugation. The resulting deposit was washed five times with PBS to remove any free APBA and Hyd. Finally, the prepared composite was dispersed in 1.0 mL PBS and stored in a refrigerator at 4 °C for further use. 2.4. Fabrication of immunosensor probe The construction of the sensor is shown in Scheme 1. A layer of the polymer of the TTBA monomer was formed on the GCE/AuNPs surface through electropolymerization of 1.0 mM TTBA monomer
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Scheme 1. Schematic representation of the immunosensor fabrication and detection principle.
solution, followed by immobilization of AntiP-gp on the surface. (details in Supporting material). 2.5. Cancer cells sample preparation The P-gp overexpressed uterine sarcoma MDRCC lines (ATCCs Number: CRL-1977™ and ATCCs Number: CRL-2274™) were obtained from the American Type Culture Collection (ATCC), Manassas, USA. Control experiments and interference studies were performed using SKBr-3, HeLa, OSE, HEK-293 cell lines which were obtained from Korean cell line bank. The cells were grown at 37 °C in 5% CO2 atmosphere in the appropriate medium supplemented with a 10% heat-inactivated FBS, 100 units/mL of penicillin and streptomycin. The MDRCC lines were cultured in McCoy’s 5A medium and maintained with 4.0 μg/mL adriamycin. Before each experiment, cells were suspended in autoclaved PBS and counted using a disposable C-Chip (South Korea) under an optical microscope (Olympus). 2.6. Detection of MDR cancer cells with the sensor probe The GCE/AuNPs/pTTBA/AntiP-gp sensor probe was incubated with enumerated MDRCC (CRL-1977™) for appropriate time (30 min) and then it was washed with the same buffer to remove the unbound cells. Next, the GCE/AuNPs/pTTBA/AntiP-gp/MDRCC electrode was incubated with APBA-MWCNT-Hyd for 30 min to form GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNT-Hyd probe followed by washing with the same buffer solution. The final GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNT-Hyd probe was tested by cycling the potential between þ0.6 and 0.7 V at a scan rate of 50.0 mV/s. The chronoamperometric experiment was carried out by applying a potential of 0.45 V vs. Ag/AgCl at the GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNT-Hyd probe into a deoxygenated 0.1 M PBS containing 4.0 mM of H2O2 (optimized) to obtained the catalytic response.
3. Results and discussion 3.1. Characterization and morphology of electrode surface Foremost, the AuNPs were electrodeposited onto the GCE using the potential step method and confirmed AuNPs deposition
employing linear sweep voltammetry (LSV). During the electrodeposition of AuNPs onto bare GCE, the peak currents increased as the number of sweep increased indicating the deposition of AuNPs and increase in the conductivity of GCE/AuNPs surface (figure not shown). The results obtained for the electrochemical deposition of AuNPs are in agreement with our previously reported results (Chandra et al., 2011b). After that, pTTBA film was formed on the GCE/AuNPs surface by electropolymerization from a 1.0 mM TTBA monomer solution. The monomer oxidation peak appeared at þ1.2 V during the anodic scan from 0.0 to þ1.4 V in a 1.0 mM monomer solution. There was a distinct reduction peak at þ0.8 V in the reverse cathodic scan from þ1.4 V, corresponding to the reduction of the oxidized polymer film formed on the GCE/AuNPs surface. These results related to nanocomposite preparation were in agreement with our previously reported results (Chandra et al., 2011a; Koh et al., 2011). To confirm the fabrication of the sensor probe, the modification steps were characterized using SEM, AFM, XPS, QCM, and EIS. Supplementary material, Fig. S1(A) shows SEM images of the (a) GCE/AuNPs, (b) GCE/pTTBA, and (c) GCE/AuNPs/pTTBA electrodes. The SEM image of the GCE/AuNPs layer shows the existence of the AuNPs of particle sizes of about 20.0 71.5 nm (Fig. S1(A(a))). The morphology of the GCE/AuNPs/pTTBA surface in the SEM image shows film of pTTBA over the deposited AuNPs (Fig. S1 (A(c))). The TEM image also confirmed the AuNPs size to be 20.07 1.5 nm (Fig. S1(B) in Supplementary material). The morphologies of the polymer layers after electropolymerization were also observed by AFM in tapping mode. The (a) HOPG/AuNPs, (b) HOPG/pTTBA, and (c) HOPG/AuNPs/pTTBA layers were formed through the electropolymerization of the monomer coated on the highly oriented pyrolytic graphite (HOPG) electrode surface (Fig. 1 (A)). The polymer AuNPs-coated HOPG electrode surface exhibits a homogeneous composition of small particles for the AuNPs/polymer films indicating the successful preparation of nanocomposite. The particle size of AuNPs/pTTBA film was determined to be 70.5 715.5 nm. The differences in the root-mean-square roughness of the surfaces among the AuNPs, pTTBA, and AuNPs/pTTBA films (3.56, 1.61, and 2.67 nm, respectively) were small. Next, we characterized the electrode surface using XPS. All XPS spectra were taken after 50 s of Ar ion gas etching and calibrated with a C1s peak at 284.6 eV as an internal standard. Fig. 1(B) shows the XPS spectra for the surfaces of (i) GCE/AuNPs/pTTBA, and (ii) GCE/ AuNPs/pTTBA/AntiP-gp. The C1s spectrum for pTTBA exhibits two
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Fig. 1. (A) Tapping mode AFM images of the (a) HOPG/AuNPs, (b) HOPG/pTTBA, and (c) HOPG/AuNPs/pTTBA films; images size is 500.0 nm 500.0 nm. (B) XPS spectra of (a) C1s, (b) S2p, and (c) N1s for (i) GCE/AuNPs/pTTBA and (ii) GCE/AuNPs/pTTBA/AntiP-gp. (C) Frequency changes during immobilization of AntiP-gp onto the Au-coated/ pTTBA chip. (D) Nyquist plots obtained for bare GCE (black), GCE/AuNPs (red), GCE/pTTBA (magenta), and GCE/AuNPs/pTTBA (blue line) electrodes in a 0.1 M PBS (pH 7.4), containing 4.0 mM Fe[(CN)6]3 /4 /0.3 M NaClO4. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
peaks at 284.6 and 286.5 eV that correspond to the C ¼C, C–C, C–H, or C–S bonds, and a C ¼O, C–O bonds, respectively (Fig. 1(B(a))). After immobilization of the antibody (GCE/AuNPs/pTTBA/AntiPgp), a new C–N bond appeared at 285.1 eV due to the bond formation between –COOH groups of pTTBA and –NH2 groups of AntiP-gp, the peak due to C ¼O bond was shifted toward the positive energy at 286.7 eV. The polymer coated film shows an S2p peak at 163.7 eV corresponding to the C–S bond, which is due to the sulfur component of pTTBA (Fig. 1(B (b))), while the peak at 163.7 eV, however, was not observed for GCE/AuNPs surface (not
shown). After the immobilization of AntiP-gp, the peaks at 399.7 and 399.1 eV appeared in the N1s spectrum corresponding to the C–N bond formation due to the covalent bond formation between –NH2 groups of AntiP-gp and –COOH groups of pTTBA indicating successful AntiP-gp immobilization (Fig. 1(B(c))). No peak at 399.7 and 399.1 eV, however, was observed for GCE/AuNPs/pTTBA surface which clearly indicates that AntiP-gp has been successfully immobilized on the nanocomposite. Further studies were performed using quartz crystal microbalance (QCM) to confirm the immobilization and estimation of the amount of AntiP-gp
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immobilized on the polymer film based on the frequency change (Fig. 1(C)). In the case of AntiP-gp immobilization, a decrease in the frequency reaches a complete steady state after about 95.0 min with an overall frequency change (Δf) of 287 Hz, corresponding to a mass change (Δm) of 315.8 712.2 ng based on a previously defined equation (Lee et al., 2010). These results reconfirm the successfully immobilization of AntiP-gp onto the pTTBA layer. Impedance spectrometry was employed to investigate the characteristics of probe layers at each modification step. Fig. 1 (D) shows Nyquist plots obtained for bare GCE, GCE/AuNPs, GCE/ pTTBA, and GCE/AuNPs/pTTBA electrodes in a 0.1 M PBS (pH 7.4), containing 4.0 mM Fe[(CN)6]3 /4 /0.3 M NaClO4. Values for Rs, Rp1, Rp2, CPE1, and CPE2 were obtained by fitting the experimental data to the equivalent circuit using the Zview2 impedance software. In the equivalent circuit, Rs represents the solution resistance, Rp1 and Rp2 represent the polarization resistances, W represents the Warburg element, and CPE1 and CPE2 are the constant-phase elements. The Rp values were determined from the cross points between the Zre axes and the extrapolation of the curve, which were obtained from a Nyquist plot of the impedance spectroscopy. Rp values for bare GCE (636.2), GCE/AuNPs (230.9), GCE/pTTBA (10006.7), and GCE/AuNPs/pTTBA (1113.1 Ω) electrodes were determined and the values decreased about nine times and the conductivity of the AuNP deposited-polymer electrode increased because of the electrodeposition of the AuNPs. 3.2. Detection of MDR cancer cells The GCE/AuNPs/pTTBA/AntiP-gp sensor was incubated with 5000 MDRCC/mL (CRL-1977™) for 30 min. Next, the GCE/AuNPs/ pTTBA/AntiP-gp/MDRCC probe was incubated with APBA-MWCNTHyd conjugate for 30 min followed by washing three times in the same buffer solution. Then, cyclic voltammetry (CV) was recorded by cycling the potential between þ0.6 and 0.7 V at a scan rate of 50.0 mV/s in deoxygenated 0.1 M PBS. In this case, a redox peak at 50/ 90 mV was observed due to the electrochemical behavior of Hyd itself indicating that APBA-MWCNT-Hyd interacted with the GCE/AuNPs/pTTBA/AntiP-gp/MDRCC electrode successfully (Fig. 2 (A), red curve). The redox peak at 50/ 90 mV was directly proportional to the scan rate between 10.0 and 60.0 mV/s indicating that the electrode reaction involved a surface-confined process of Hyd. In two separate controls we performed the similar experiment as mentioned above except Hyd and APBA. In both the cases no redox current was observed (blue and black curve). This is due to absence of Hyd, an electrochemically active group and APBA, the recognition molecule on the sensor surface. The results from these control experiments clearly suggest that APBA and Hyd both are required to obtain an electrochemical signal to detect the MDRCC. The ability of GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNTHyd probe to catalyze the reduction of H2O2 was further investigated. The sensor probe was dipped in a deoxygenated 0.1 M PBS containing 4.0 mM H2O2 and the CVs was recorded. A very distinct reduction peak at 400 mV was observed due to the reduction of H2O2 by Hyd attached on the sensor probe (Fig. 2(B)). No reduction peak at 400 mV, however, was observed when control experiments were performed without APBA and Hyd. Next, we studied the H2O2 catalytic current with increasing MDRCC to establish that the current arose due to the immunoreaction. Fig. 2 (B) shows the CVs recorded with the increasing number of MDRCC, where the current response increases with increase in the MDRCC from 1000 to 3500 cells/mL. In this figure, a represents blank and b (1000), c (1500), d (2000), e (2500), f (3000), and g (3500) represents MDRCC/mL. A linear plot is obtained (inset) with a linear regression equation expressed as follows: ΔI (μA)¼ 2.43þ0.002 [MDRCC], and has a correlation coefficient of 0.997. These results clearly indicate that the developed biosensor can accurately detect
Fig. 2. (A) CV response at GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA (black line) GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/MWCNT-Hyd (blue line), GCE/AuNPs/pTTBA/ AntiP-gp/ MDRCC/APBA-MWCNT-Hyd (red line) at the scan rate of 50.0 mV/s, 5000 MDRCC/mL were used in all these experiments. (B) CVs recorded for the catalytic reduction of 4.0 mM H2O2 at GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNTHyd sensor with increasing number of MDRCC/mL [a represents blank, b (1000), c (1500), d (2000), e (2500), f (3000), and g (3500) MDRCC/mL]. Inset shows the calibration plot for increasing number of MDRCC/mL. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
MDRCC. Next, we optimized the experimental parameters of the sensor probe to obtain the detection limit of MDRCC using chronoamperometry (see details in Fig. S2 in Supplementary material). 3.3. Analytical performance of the sensor Analytical performance of the GCE/AuNPs/pTTBA/AntiP-gp sensor probe was tested for the increasing numbers of MDRCC (CRL-1977™ cells) followed by binding with APBA-MWCNT-Hyd conjugate and reaction with H2O2. Under the optimized conditions chronoamperometry was performed for GCE/AuNPs/pTTBA/AntiP-gp/ MDRCC/APBA-MWCNT-Hyd probe at 0.45 V vs. Ag/AgCl (optimized) in deoxygenated PBS containing 4.0 mM H2O2. The increase in the catalytic signal due to H2O2 reduction was obtained in chronoamperometry which was directly proportional to the amount of MDRCC captured by the sensor. Fig. 3(A) shows the chronoamperogram with gradual increased in the current response with increasing
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Fig. 3. (A) Amperometic responses for MDRCC detection using the biosensor with increasing number of MDRCC/mL [a represents blank, b (50), c (100), d (500), e (1000), f (10,000), g (25,000), h (50,000), i (75,000), and j (100,000) MDRCC/mL]. (B) Calibration plot based on the signal obtained from the amperogram (inset shows the magnified values between 50 and 10,000 MDRCC/mL). All experiments were performed in deoxygenated 0.1 M PBS containing 4.0 mM H2O2 (pH 7.4), applied potential: 0.45 V vs. Ag/AgCl.
number of MDRCC. In this figure, a represents blank and b (50), c (100), d (500), e (1000), f (10,000), g (25,000), h (50,000), i (75,000), and j (100,000) represents MDRCC/mL. Based on the obtained chronoamperograms a calibration curve was plotted as shown in Fig. 3(B). The calibration curve shows a linear response with the MDRCC ranging from 50 to 100,000 cells/mL. The linear regression equation is expressed as follows: ΔI¼0.9351 (70.1181)þ0.0025 (70.0000067) [MDRCC] with the correlation coefficient of 0.998. The detection limit of MDRCC was determined to be 2372 cells/mL (RSDo4.4%) based on the standard deviation of five repeated measurements of the blank (95% confidence level, n¼ 5). The obtained detection limit is 4 times lower compared to the recently reported MDRCC impedimetric biosensor based on gold nanoparticles/polyaniline nanofibers (Zhang et al., 2014) and also to other recently reported chemiluminescence (Chen et al., 2014) and QCM (Shaolian et al., 2014) biosensor for cancer cell detection. We also performed same experiment with another P-gp over expressed cell line obtained from ATCC, USA (CRL-2274™). The current response increased with increase in the CRL-2274™ cells, however, the signals in this case was 17% lower compared to the CRL-1977™ cell type (data not
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Fig. 4. (A) Histogram showing the selective detection of MDRCC using biosensor at 30,000 MDRCC/mL concentration. (B) Microphotograph of MDRCC captured by ITO/ pTTBA/AntiP-gp probe ((a–c)) and (d) ITO/pTTBA chip treated with MDRCC (no cells were observed after washing the electrode).
shown). This is possibly due to the less concentration of the P-gp expressed on the cell membrane of CRL-2274™cells. 3.4. Selectivity, reproducibility, stability, and real sample analysis Control experiments were performed to validate the selectivity of the developed nanobiosensor. For this purpose, SKBr-3, HeLa, OSE, and HEK-293 cells were examined under the similar experimental conditions. The selectivity test was performed at 1000, 30,000, 75,000, and 100,000 cells/mL concentrations. Fig. 4 (A) shows the current response at 30,000 cells/mL concentration, where negligible signal is observed for tested cells. This was because no immunoreaction occurred between AntiP-gp and cell surface P-gp which was not present (or negligible) on these cells. We also performed control experiments to ensure the selectivity of the nanobiosensor by testing the compounds that are majorly present in the real sample matrix such as; albumin, fibrinogens, glucose. Interestingly, no current response was observed for these compounds indicating that the develop nanobiosensor is highly selective towards MDRCC detection and no interference or false
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positive signal is obtained. The reproducibility of the analysis was performed which showed the RSD o3.5% (n ¼5) and electrode-toelectrode RSD was o 3.2% even when the same preparation conditions were applied. This minor variation was possibly due to the difference in the modified surface and/or slight variations in the optimized conditions. We also studied the stability of the sensor probe with respect to time. The immunosensor retained almost 96% of its sensitivity for 6 weeks. The response decreased about 13% after 6 weeks and then gradually decreased with time. Thus, the sensor was stable for 6 weeks. The good stability was probably caused by the strong covalent attachment of the antibody on the stable conducting polymer–nanoparticle composite. The biomedical application of the nanobiosensor was investigated by detecting MDRCC in a human serum samples in the similar dynamic range (50–100,000 MDRCC/mL) as done in the buffer solution. Serum has been widely used as a model sample to demonstrate the biomedical application of biosensors (Chandra et al., 2013; Colin et al., 2011; Joshua et al., 2007). Thus, we used serum to examine the real clinical value of the developed sensor. MDRCC were dispersed into five times diluted serum samples and incubated with the GCE/AuNPs/pTTBA/AntiP-gp sensor probe. After washing the sensor probe was treated with the APBAMWCNT-Hyd conjugate, and reacted with H2O2 as described earlier. The current response increased with increase in the MDRCC number indicating that the developed biosensor can effectively detect MDRCC in serum samples, hence valuable in the biomedical analysis. The linear regression equation for the calibration plot in the serum sample is expressed as follows: ΔI ¼0.881 þ0.0023 [MDRCC] with the correlation coefficient of 0.998 indicating that this sensor is capable to detect MDRCC from the complex biological sample matrix effectively. The detection limit of MDRCC in serum sample was determined to be 2872 cells/mL (RSD o4.8%) based on the standard deviation of five repeated measurements of the blank (95% confidence level, n¼ 5). The detection limit in serum sample was 5 cells/mL higher (less sensitive) compared to the detection limit in a buffer solution (23 cells/mL). This minor variation in the detection limit was possibly due to negligible matrix effect due to the serum components. We also examined the applicability of the biosensor to detect MDRCC in the mixed cell sample. For this purpose, enumerated MDRCC were mixed with SKBr-3, HeLa, OSE, and HEK-293 cell and detected by the biosensor probe. Interestingly, the sensitivity of the detection in this case was 97% (n ¼5) compared when MDRCC was tested alone. This result clearly indicates that the developed biosensor can selectively detect MDRCC in the presence of other cancerous and noncancerous cells effectively. 3.5. Bioimagining and comparison with conventional reporter antibody based assay The capturing of the MDRCC by the AntiP-gp probe was investigated by bioimagining MDRCC incubated ITO/pTTBA/AntiP-gp surface using optical microscope. The ITO/pTTBA/AntiP-gp sensor was incubated with the MDRCC and after washing it was observed under microscope. Fig. 4(B) shows the microscopic images of ITO/ pTTBA/AntiP-gp sensor after treating it with different concentrations of MDRCC. With increase in the MDRCC concentrations between 102 to 103 cells/mL, the number of cells captured by the sensor was increased as shown in Fig. 4(B(a–c)). These results clearly indicate that ITO/pTTBA/AntiP-gp sensor is able to detect MDRCC. No MDRCC, however, were observed when Anti-gp was not immobilized on the ITO/pTTBA electrode indicating that AntiP-gp is essential to detect MDRCC (Fig. 4(B(d))). We also compared this method with the conventional reporter antibody based immunoassay. In this experiment, a reporter AntiP-gp-MWCNT-Hyd conjugate was prepared through the carbodiimide coupling
Fig. 5. Histogram showing the comparative response of the APBA-MWCNT-Hyd and reporter AntiP-gp-MWCNT-Hyd conjugate. The values on the y-axis shown in the histograms are based on the amperometric signals.
reaction following our earlier work (Zhu et al., 2010). This conjugate was tested by the following same experimental steps as described in the previous section. Fig. 5 shows the comparative response of the APBA-MWCNT-Hyd and the reporter AntiP-gpMWCNT-Hyd conjugate at 25,000, 50,000, and 100,000 MDRCC/mL. A much higher signal was observed for APBA-MWCNTHyd conjugate. The detection limit of MDRCC using reporter AntiPgp-MWCNT-Hyd conjugate was 15878 cells/mL based on the standard deviation of five repeated measurements of the blank (95% confidence level, n ¼5). This detection limit is 7 times higher i.e. less sensitive compared to the detection limit obtained using APBA-MWCNT-Hyd conjugate (237 2 cells/mL). This is due to the less immunoreaction between the reporter AntiP-gpMWCNT-Hyd conjugate and the MDRCC captured by the sensor probe. This less immunoreaction is due to the low concentrations of P-gp on the MDRCC cell membrane. It has already been reported that variation in the concentration of cell surface antigen (e.g.: P-gp in our case) significantly affects the antibody binding (Langmuir et al., 1991; Velders et al., 1998) which may directly influence the sensitivity of the biosensor. In case of APBA-MWCNTHyd conjugate low P-gp concentration does not affect the sensitivity of the biosensor because the target molecule for this conjugate is glycan which is expressed greatly onto the cancer cell surface (Zhang et al., 2010).
4. Conclusions A highly sensitive and selective amperometric nanobiosensor has been successfully designed for the detection of multidrug resistant cancer cells (MDRCC) in the biological matrix. The selectivity and sensitivity of MDRCC detection were achieved due to the monoclonal Pg-antibody covalently immobilized onto the conduction polymer-gold nanoparticle composite and also due the APBA-MWCNT-Hyd conjugate. The developed nanobiosensor is able to detect the MDRCC in mixed cell sample and in the presence of other chemical molecules. The detection limit of nanobiosensor is 23 72 MDRCC/mL, which is four times more sensitive than the most recent MDR cancer cell biosensor. The developed nanobiosensor is 7 times more sensitive compared to the conventional reporter antibody based immunosensor which gives a new direction towards sensitive detection of cancer cells. This method
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proposed herein is a generic method and can be easily adopted for other type of cancer cells simply by changing the detector antibody. The developed sensor could be a sensitive diagnostic method for the detection of drug resistant cancer cell in the biological fluids of cancer patients which will help clinicians to develop their alternate therapeutic strategies. The study clearly demonstrates highly sensitive detection of cancer cells even though the cell surface biomarkers is expressed/present in low concentrations using a nanobiosensor. In future, the developed nanobiosensor prototype can be translated into a miniaturized kit based assay for the point-of-care medical applications.
Acknowledgment This work was supported by the National Research Foundation of Korea (NRF) Grant funded by the Korea government (MSIP) (No. 20100029128).
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.2015.03.069.
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Biosensors and Bioelectronics journal homepage: www.elsevier.com/locate/bios
Ultrasensitive detection of drug resistant cancer cells in biological matrixes using an amperometric nanobiosensor Pranjal Chandra a,b, Hui-Bog Noh a, Ramjee Pallela a, Yoon-Bo Shim a,n a b
Department of Chemistry and Institute of Biophysio Sensor Technology (IBST), Pusan National University, Busan 609-735, South Korea Amity Institute of Biotechnology, Amity University Uttar Pradesh, Noida 201303, U.P., India
art ic l e i nf o
a b s t r a c t
Article history: Received 5 February 2015 Received in revised form 10 March 2015 Accepted 26 March 2015 Available online 27 March 2015
Multidrug resistance (MDR) is a key issue in the failure of cancer chemotherapy and its detection will be helpful to develop suitable therapeutic strategies for cancer patients and overcome the death rates. In this direction, we designed a new amperometric sensor (a medical device prototype) to detect drug resistant cancer cells by sensing “Permeability glycoprotein (P-gp)”. The sensor probe is fabricated by immobilizing monoclonal P-gp antibody on the gold nanoparticles (AuNPs) conducting polymer composite. The detection relies on a sandwich-type approach using a bioconjugate, where the aminophenyl boronic acid (APBA) served as a recognition molecule which binds with the cell surface glycans and hydrazine (Hyd) served as an electrocatalyst for the reduction of H2O2 which are attached on multi-wall carbon nanotube (MWCNT) (APBA-MWCNT-Hyd). A linear range for the cancer cell detection is obtained between 50 and 100,000 cells/mL with the detection limit of 237 2 cells/mL. The proposed immunosensor is successfully applied to detect MDR cancer cells (MDRCC) in serum and mixed cell samples. Interferences by drug sensitive (SKBr-3 and HeLa), noncancerous cells (HEK-293 and OSE), and other chemical molecules present in the real sample matrix are examined. The sensitivity of the proposed immunosensor is excellent compared with the conventional reporter antibody based assay. & 2015 Elsevier B.V. All rights reserved.
Keywords: Nanobiosensor Drug resistant cancer cells Electrochemistry In vitro test Nanomaterials
1. Introduction Multidrug resistance (MDR) is a major factor in the failure of cancer chemotherapy. Several mechanisms have been convinced to play vital roles in the development of the MDR in cancer cells (Krishna and Mayer, 2000). Among them, a major form is mediated by a cell membrane transporter “Permeability glycoprotein, (P-gp)” which is encoded by the MDR1 gene in human cells. It includes two nucleotide-binding and two membrane-spanning domains, acts as energy dependent pump that decreases intracellular drug concentrations lower than the effective therapeutic levels (Bellamy, 1996; Gottesman et al., 2002). This phenomenon leads to the development of drug resistance in cancer cells. The early detection of MDR in cancer cells can lead to the adequate chemotherapy of cancer patients and overcome the mortality. Thus, it is extremely important and clinically significant to detect MDRCC at the early stage. This will help the clinicians to develop alternate therapeutic strategies for cancer patients. Until now, several biological methods have been developed to detect P-gp expression for the diagnosis of MDR in cancer such as; n
Corresponding author. Fax: þ82 51 514 2122. E-mail address: [email protected] (Y.-B. Shim).
http://dx.doi.org/10.1016/j.bios.2015.03.069 0956-5663/& 2015 Elsevier B.V. All rights reserved.
polymerase chain reaction (Murphy et al., 1990), immunohistochemistry (Chan et al., 1990), flow cytometry (Ludescher et al., 1992), and microarray (Gillet et al., 2004). Although these methods can be used to detect MDRCC, they are less sensitive, require highly trained professionals, and have no ability to be miniaturized for the point-of-care applications. Thus, few biosensor-based diagnostic methods have also been reported to detect MDRCC. These methods, however, are indirect, nonselective, and less sensitive (Du et al., 2005; Zhang et al., 2011, 2014). Thus, the development of a sensitive, robust, and alternate method for the detection of MDRCC directly in the biological fluids is desirable and has great clinical importance. To achieve a highly sensitive detection method for MDRCC, a sandwich type-electrochemical immunosensor composed of P-gp monoclonal antibody (AntiP-gp) as a detection probe coupled with a ligand tagged-non-enzymatic catalyst as a reporter probe is worthy to attempt due to their selective and highly sensitive characteristics, respectively. The stable immobilization of biomaterials including antibodies on the sensor probe is extremely critical in the biosensor fabrication (Rahman et al., 2008). Electrochemical biosensors composed of conducting polymers-AuNPs composite are considered to be highly stable and ultrasensitive because biomolecules can be covalently immobilized on the polymer backbone possessing –COOH or –NH2 groups (Chandra et al., 2011a, 2013; Koh et al.,
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2011; Kim and Shim, 2013; Lee et al., 2010). Compare to the reporter antibody based conventional sandwiched type electrochemical immunosensor (Malhotra et al. 2012; Wu et al., 2015; Zhu et al. 2014), other ligands are worthy to attempt for the in vitro detection of cancer cells. This is important because it has been reported that the concentration of antigen (target molecule) on the cancer cell surface is low (Beck et al., 1996). This may result in low antigen–antibody binding and consequently poor detection of cancer cells. Thus, it is worthy to target molecules that are present on the cancer cell surface in abundance. Recent studies have clearly indicated existence of high concentration of glycans onto the cancer cell membrane (Dube and Bertozzi, 2005; Zhang et al., 2010) due to which numerous diol-groups are present on the cell surface. Arising from the unique capacity of boronic acids to form boronic esters with these diols (Das et al., 2011), it would be interesting to attempt boronic acid as an alternative recognition molecule (reporter probe) to detect cancer cell surface glycans. To show the signal in an electrochemical immunosensor, an electrochemical indicator is needed, such as a bioconjugate composed of enzyme or nonenzymatic catalyst. Compared to expensive and easily denaturable enzyme, Hyd attached on a bioconjugate can be used due to its small size, stability, and high catalytic activity towards hydrogen peroxide (H2O2) reduction (Zhu et al., 2013). Thus, in this work we developed a sandwich immunoassay approach having a chemo-nano-conjugate composed of APBA, MWCNT, and Hyd instead of enzyme linked reporter antibody. In this conjugate, APBA and Hyd are used a reporter molecule and electrocatalyst for signal generation, respectively. In the present study, a novel amperometric immunosensor for the detection of MDRCC has been tried to develop for the first time through the detection of P-gp. The immunosensor probe is fabricated by forming covalent bonds between AntiP-gp and carboxylic acid group-functionalized conducting polymer layer on AuNPs deposited electrode surface. The nanocomposite was characterized by atomic force microscopy (AFM), X-ray photoelectron spectroscopy (XPS), electrochemical impedance spectroscopy (EIS), scanning electron microscopy (SEM), and the AntiP-gp immobilization was confirmed by quartz crystal microbalance (QCM) and XPS. The sandwich immunosensing approach was adopted where APBAMWCNT-Hyd was reacted with the MDRCC captured by the GCE/ AuNPs/pTTBA/AntiP-gp probe. The immunoreaction was monitored in term of catalytic activity of Hyd towards H2O2 reduction. The experimental parameters were optimized and the detection limit of the MDRCC was determined. Direct detection of MDRCC in serum and in the mixed cell samples were performed to evaluate the real clinical value of sensor. The selectivity of the biosensor was also examined toward various non target cells, and chemicals present in the real sample matrix. The developed method was also compared with the conventional reporter antibody based method.
2. Material and methods 2.1. Materials 2,2′:5′, 2″-terthiophene-3′(p-benzoic acid) (TTBA) was synthesized through the Paal–Knorr pyrrole condensation reaction (Koh et al., 2011). Tetrabutylammonium perchlorate (TBAP, electrochemical grade) was purchased from Fluka (USA) and purified according to a general method, followed by drying under vacuum at 1.33 10 3 Pa (Noh et al., 2012). 1-ethyl-3-(3-(dimethylamino)propyl) carbodiimide (EDC), N-Hydroxysuccinimide (NHS), dichloromethane (99.8%, anhydrous), trisodium citrate, sodium tetrahydridoborate, HAuCl4 3H2O, bovine serum albumin (BSA), and Hydrazine sulfate were purchased from sigma Aldrich (USA).
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Monoclonal p-GP antibody, amino-phenylboronic acid, indium tin oxide (ITO) glass, and H2O2 (33%) were purchased from sigma Aldrich (USA). MWCNTs (4–6 nm diameter, 95%) were obtained from Iljin Nanotech (South Korea). phosphate buffer saline solutions (PBS) were prepared with 0.1 M of disodium Hydrogen phosphate, 0.1 M of sodium dihydrogenphosphate, and 0.9% sodium chloride. cell culture medium, fetal bovine serum (FBS), trypsin-EDTA, penicillin/streptomycin, hank’s balance salt (HBS) solution, were purchased from Sigma-Aldrich (USA). All other chemicals were of extra pure analytical grade and used without further purification. all aqueous solutions were prepared in ultrapure water, which was obtained from a Milli-Q water purifying system (18 M Ω cm). 2.2. Apparatus All electrochemical experiments were performed in a three electrode cell. The modified glassy carbon electrode (GCE) (dia. 3.0 mm), Ag/AgCl (in saturated KCl), and a platinum (Pt) wire were used as working, reference, and counter electrodes, respectively. Voltammograms and chronoamperograms were recorded using a potentiostat/galvanostat, Kosentech, model KST P-2 (South Korea). A Multimode AFM device (Veeco Metrology) equipped with a Nanoscope IV controller (Veeco) was used at ambient conditions to get the images. The QCM experiment was performed using a SEIKO EG&G model QCA 917 and a PAR model 263A potentiostat/ galvanostat (USA). An Au-coated working electrode (area: 0.196 cm2; 9 MHz; AT-cut quartz crystal) was used for the QCM experiment. The impedance spectra were obtained using a EG&G Princeton Applied Research PARSTAT2263 at an open circuit voltage from 100.0 kHz down to 100.0 mHz and a sampling rate of five points per decade (AC amplitude: 10.0 mV). XPS was performed using a VG Scientific XPSLAB 250 XPS spectrometer and a monochromated Al Kα source with charge compensation at KBSI (Busan). The SEM images were obtained using a Cambridge Stereoscan 240. A JEOL JEM-2010 electron microscope (Jeol High-Tech Co., Japan) with an acceleration voltage of 200 kV was used to obtain TEM images. 2.3. Preparation of APBA-MWCNT-Hyd conjugate Firstly, MWCNTs were functionalized according to previously reported methods (Goldman and Lellouche, 2010; Piran et al., 2009). Briefly, 100 mg of MWCNT was treated with mixture of concentrated 12.0 M HNO3 and 36.0 M H2SO4 (90 °C, 2.0 h) followed by multiple rinsing with deionized water until no acid was detected. After drying at 80 °C under vacuum overnight, a black powder was obtained. Next, the treated MWCNT (3.0 mg) was dispersed into 1.0 mL PBS (pH 7.0) containing 10.0 mM EDC/NHS solutions and incubated at room temperature for 6 h to activate the –COOH on the MWCNTs. The resulting mixture was separated by centrifugation and the precipitate was washed three times. Meanwhile, an optimized 1.0 mg/mL hydrazine sulfate solution and 10.0 mM APBA was prepared in PBS. Subsequently, the activated MWCNTs were mixed with hydrazine sulfate and APBA solution. The mixture was stored in a refrigerator overnight at 4 °C followed by centrifugation. The resulting deposit was washed five times with PBS to remove any free APBA and Hyd. Finally, the prepared composite was dispersed in 1.0 mL PBS and stored in a refrigerator at 4 °C for further use. 2.4. Fabrication of immunosensor probe The construction of the sensor is shown in Scheme 1. A layer of the polymer of the TTBA monomer was formed on the GCE/AuNPs surface through electropolymerization of 1.0 mM TTBA monomer
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Scheme 1. Schematic representation of the immunosensor fabrication and detection principle.
solution, followed by immobilization of AntiP-gp on the surface. (details in Supporting material). 2.5. Cancer cells sample preparation The P-gp overexpressed uterine sarcoma MDRCC lines (ATCCs Number: CRL-1977™ and ATCCs Number: CRL-2274™) were obtained from the American Type Culture Collection (ATCC), Manassas, USA. Control experiments and interference studies were performed using SKBr-3, HeLa, OSE, HEK-293 cell lines which were obtained from Korean cell line bank. The cells were grown at 37 °C in 5% CO2 atmosphere in the appropriate medium supplemented with a 10% heat-inactivated FBS, 100 units/mL of penicillin and streptomycin. The MDRCC lines were cultured in McCoy’s 5A medium and maintained with 4.0 μg/mL adriamycin. Before each experiment, cells were suspended in autoclaved PBS and counted using a disposable C-Chip (South Korea) under an optical microscope (Olympus). 2.6. Detection of MDR cancer cells with the sensor probe The GCE/AuNPs/pTTBA/AntiP-gp sensor probe was incubated with enumerated MDRCC (CRL-1977™) for appropriate time (30 min) and then it was washed with the same buffer to remove the unbound cells. Next, the GCE/AuNPs/pTTBA/AntiP-gp/MDRCC electrode was incubated with APBA-MWCNT-Hyd for 30 min to form GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNT-Hyd probe followed by washing with the same buffer solution. The final GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNT-Hyd probe was tested by cycling the potential between þ0.6 and 0.7 V at a scan rate of 50.0 mV/s. The chronoamperometric experiment was carried out by applying a potential of 0.45 V vs. Ag/AgCl at the GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNT-Hyd probe into a deoxygenated 0.1 M PBS containing 4.0 mM of H2O2 (optimized) to obtained the catalytic response.
3. Results and discussion 3.1. Characterization and morphology of electrode surface Foremost, the AuNPs were electrodeposited onto the GCE using the potential step method and confirmed AuNPs deposition
employing linear sweep voltammetry (LSV). During the electrodeposition of AuNPs onto bare GCE, the peak currents increased as the number of sweep increased indicating the deposition of AuNPs and increase in the conductivity of GCE/AuNPs surface (figure not shown). The results obtained for the electrochemical deposition of AuNPs are in agreement with our previously reported results (Chandra et al., 2011b). After that, pTTBA film was formed on the GCE/AuNPs surface by electropolymerization from a 1.0 mM TTBA monomer solution. The monomer oxidation peak appeared at þ1.2 V during the anodic scan from 0.0 to þ1.4 V in a 1.0 mM monomer solution. There was a distinct reduction peak at þ0.8 V in the reverse cathodic scan from þ1.4 V, corresponding to the reduction of the oxidized polymer film formed on the GCE/AuNPs surface. These results related to nanocomposite preparation were in agreement with our previously reported results (Chandra et al., 2011a; Koh et al., 2011). To confirm the fabrication of the sensor probe, the modification steps were characterized using SEM, AFM, XPS, QCM, and EIS. Supplementary material, Fig. S1(A) shows SEM images of the (a) GCE/AuNPs, (b) GCE/pTTBA, and (c) GCE/AuNPs/pTTBA electrodes. The SEM image of the GCE/AuNPs layer shows the existence of the AuNPs of particle sizes of about 20.0 71.5 nm (Fig. S1(A(a))). The morphology of the GCE/AuNPs/pTTBA surface in the SEM image shows film of pTTBA over the deposited AuNPs (Fig. S1 (A(c))). The TEM image also confirmed the AuNPs size to be 20.07 1.5 nm (Fig. S1(B) in Supplementary material). The morphologies of the polymer layers after electropolymerization were also observed by AFM in tapping mode. The (a) HOPG/AuNPs, (b) HOPG/pTTBA, and (c) HOPG/AuNPs/pTTBA layers were formed through the electropolymerization of the monomer coated on the highly oriented pyrolytic graphite (HOPG) electrode surface (Fig. 1 (A)). The polymer AuNPs-coated HOPG electrode surface exhibits a homogeneous composition of small particles for the AuNPs/polymer films indicating the successful preparation of nanocomposite. The particle size of AuNPs/pTTBA film was determined to be 70.5 715.5 nm. The differences in the root-mean-square roughness of the surfaces among the AuNPs, pTTBA, and AuNPs/pTTBA films (3.56, 1.61, and 2.67 nm, respectively) were small. Next, we characterized the electrode surface using XPS. All XPS spectra were taken after 50 s of Ar ion gas etching and calibrated with a C1s peak at 284.6 eV as an internal standard. Fig. 1(B) shows the XPS spectra for the surfaces of (i) GCE/AuNPs/pTTBA, and (ii) GCE/ AuNPs/pTTBA/AntiP-gp. The C1s spectrum for pTTBA exhibits two
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Fig. 1. (A) Tapping mode AFM images of the (a) HOPG/AuNPs, (b) HOPG/pTTBA, and (c) HOPG/AuNPs/pTTBA films; images size is 500.0 nm 500.0 nm. (B) XPS spectra of (a) C1s, (b) S2p, and (c) N1s for (i) GCE/AuNPs/pTTBA and (ii) GCE/AuNPs/pTTBA/AntiP-gp. (C) Frequency changes during immobilization of AntiP-gp onto the Au-coated/ pTTBA chip. (D) Nyquist plots obtained for bare GCE (black), GCE/AuNPs (red), GCE/pTTBA (magenta), and GCE/AuNPs/pTTBA (blue line) electrodes in a 0.1 M PBS (pH 7.4), containing 4.0 mM Fe[(CN)6]3 /4 /0.3 M NaClO4. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
peaks at 284.6 and 286.5 eV that correspond to the C ¼C, C–C, C–H, or C–S bonds, and a C ¼O, C–O bonds, respectively (Fig. 1(B(a))). After immobilization of the antibody (GCE/AuNPs/pTTBA/AntiPgp), a new C–N bond appeared at 285.1 eV due to the bond formation between –COOH groups of pTTBA and –NH2 groups of AntiP-gp, the peak due to C ¼O bond was shifted toward the positive energy at 286.7 eV. The polymer coated film shows an S2p peak at 163.7 eV corresponding to the C–S bond, which is due to the sulfur component of pTTBA (Fig. 1(B (b))), while the peak at 163.7 eV, however, was not observed for GCE/AuNPs surface (not
shown). After the immobilization of AntiP-gp, the peaks at 399.7 and 399.1 eV appeared in the N1s spectrum corresponding to the C–N bond formation due to the covalent bond formation between –NH2 groups of AntiP-gp and –COOH groups of pTTBA indicating successful AntiP-gp immobilization (Fig. 1(B(c))). No peak at 399.7 and 399.1 eV, however, was observed for GCE/AuNPs/pTTBA surface which clearly indicates that AntiP-gp has been successfully immobilized on the nanocomposite. Further studies were performed using quartz crystal microbalance (QCM) to confirm the immobilization and estimation of the amount of AntiP-gp
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immobilized on the polymer film based on the frequency change (Fig. 1(C)). In the case of AntiP-gp immobilization, a decrease in the frequency reaches a complete steady state after about 95.0 min with an overall frequency change (Δf) of 287 Hz, corresponding to a mass change (Δm) of 315.8 712.2 ng based on a previously defined equation (Lee et al., 2010). These results reconfirm the successfully immobilization of AntiP-gp onto the pTTBA layer. Impedance spectrometry was employed to investigate the characteristics of probe layers at each modification step. Fig. 1 (D) shows Nyquist plots obtained for bare GCE, GCE/AuNPs, GCE/ pTTBA, and GCE/AuNPs/pTTBA electrodes in a 0.1 M PBS (pH 7.4), containing 4.0 mM Fe[(CN)6]3 /4 /0.3 M NaClO4. Values for Rs, Rp1, Rp2, CPE1, and CPE2 were obtained by fitting the experimental data to the equivalent circuit using the Zview2 impedance software. In the equivalent circuit, Rs represents the solution resistance, Rp1 and Rp2 represent the polarization resistances, W represents the Warburg element, and CPE1 and CPE2 are the constant-phase elements. The Rp values were determined from the cross points between the Zre axes and the extrapolation of the curve, which were obtained from a Nyquist plot of the impedance spectroscopy. Rp values for bare GCE (636.2), GCE/AuNPs (230.9), GCE/pTTBA (10006.7), and GCE/AuNPs/pTTBA (1113.1 Ω) electrodes were determined and the values decreased about nine times and the conductivity of the AuNP deposited-polymer electrode increased because of the electrodeposition of the AuNPs. 3.2. Detection of MDR cancer cells The GCE/AuNPs/pTTBA/AntiP-gp sensor was incubated with 5000 MDRCC/mL (CRL-1977™) for 30 min. Next, the GCE/AuNPs/ pTTBA/AntiP-gp/MDRCC probe was incubated with APBA-MWCNTHyd conjugate for 30 min followed by washing three times in the same buffer solution. Then, cyclic voltammetry (CV) was recorded by cycling the potential between þ0.6 and 0.7 V at a scan rate of 50.0 mV/s in deoxygenated 0.1 M PBS. In this case, a redox peak at 50/ 90 mV was observed due to the electrochemical behavior of Hyd itself indicating that APBA-MWCNT-Hyd interacted with the GCE/AuNPs/pTTBA/AntiP-gp/MDRCC electrode successfully (Fig. 2 (A), red curve). The redox peak at 50/ 90 mV was directly proportional to the scan rate between 10.0 and 60.0 mV/s indicating that the electrode reaction involved a surface-confined process of Hyd. In two separate controls we performed the similar experiment as mentioned above except Hyd and APBA. In both the cases no redox current was observed (blue and black curve). This is due to absence of Hyd, an electrochemically active group and APBA, the recognition molecule on the sensor surface. The results from these control experiments clearly suggest that APBA and Hyd both are required to obtain an electrochemical signal to detect the MDRCC. The ability of GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNTHyd probe to catalyze the reduction of H2O2 was further investigated. The sensor probe was dipped in a deoxygenated 0.1 M PBS containing 4.0 mM H2O2 and the CVs was recorded. A very distinct reduction peak at 400 mV was observed due to the reduction of H2O2 by Hyd attached on the sensor probe (Fig. 2(B)). No reduction peak at 400 mV, however, was observed when control experiments were performed without APBA and Hyd. Next, we studied the H2O2 catalytic current with increasing MDRCC to establish that the current arose due to the immunoreaction. Fig. 2 (B) shows the CVs recorded with the increasing number of MDRCC, where the current response increases with increase in the MDRCC from 1000 to 3500 cells/mL. In this figure, a represents blank and b (1000), c (1500), d (2000), e (2500), f (3000), and g (3500) represents MDRCC/mL. A linear plot is obtained (inset) with a linear regression equation expressed as follows: ΔI (μA)¼ 2.43þ0.002 [MDRCC], and has a correlation coefficient of 0.997. These results clearly indicate that the developed biosensor can accurately detect
Fig. 2. (A) CV response at GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA (black line) GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/MWCNT-Hyd (blue line), GCE/AuNPs/pTTBA/ AntiP-gp/ MDRCC/APBA-MWCNT-Hyd (red line) at the scan rate of 50.0 mV/s, 5000 MDRCC/mL were used in all these experiments. (B) CVs recorded for the catalytic reduction of 4.0 mM H2O2 at GCE/AuNPs/pTTBA/AntiP-gp/MDRCC/APBA-MWCNTHyd sensor with increasing number of MDRCC/mL [a represents blank, b (1000), c (1500), d (2000), e (2500), f (3000), and g (3500) MDRCC/mL]. Inset shows the calibration plot for increasing number of MDRCC/mL. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
MDRCC. Next, we optimized the experimental parameters of the sensor probe to obtain the detection limit of MDRCC using chronoamperometry (see details in Fig. S2 in Supplementary material). 3.3. Analytical performance of the sensor Analytical performance of the GCE/AuNPs/pTTBA/AntiP-gp sensor probe was tested for the increasing numbers of MDRCC (CRL-1977™ cells) followed by binding with APBA-MWCNT-Hyd conjugate and reaction with H2O2. Under the optimized conditions chronoamperometry was performed for GCE/AuNPs/pTTBA/AntiP-gp/ MDRCC/APBA-MWCNT-Hyd probe at 0.45 V vs. Ag/AgCl (optimized) in deoxygenated PBS containing 4.0 mM H2O2. The increase in the catalytic signal due to H2O2 reduction was obtained in chronoamperometry which was directly proportional to the amount of MDRCC captured by the sensor. Fig. 3(A) shows the chronoamperogram with gradual increased in the current response with increasing
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Fig. 3. (A) Amperometic responses for MDRCC detection using the biosensor with increasing number of MDRCC/mL [a represents blank, b (50), c (100), d (500), e (1000), f (10,000), g (25,000), h (50,000), i (75,000), and j (100,000) MDRCC/mL]. (B) Calibration plot based on the signal obtained from the amperogram (inset shows the magnified values between 50 and 10,000 MDRCC/mL). All experiments were performed in deoxygenated 0.1 M PBS containing 4.0 mM H2O2 (pH 7.4), applied potential: 0.45 V vs. Ag/AgCl.
number of MDRCC. In this figure, a represents blank and b (50), c (100), d (500), e (1000), f (10,000), g (25,000), h (50,000), i (75,000), and j (100,000) represents MDRCC/mL. Based on the obtained chronoamperograms a calibration curve was plotted as shown in Fig. 3(B). The calibration curve shows a linear response with the MDRCC ranging from 50 to 100,000 cells/mL. The linear regression equation is expressed as follows: ΔI¼0.9351 (70.1181)þ0.0025 (70.0000067) [MDRCC] with the correlation coefficient of 0.998. The detection limit of MDRCC was determined to be 2372 cells/mL (RSDo4.4%) based on the standard deviation of five repeated measurements of the blank (95% confidence level, n¼ 5). The obtained detection limit is 4 times lower compared to the recently reported MDRCC impedimetric biosensor based on gold nanoparticles/polyaniline nanofibers (Zhang et al., 2014) and also to other recently reported chemiluminescence (Chen et al., 2014) and QCM (Shaolian et al., 2014) biosensor for cancer cell detection. We also performed same experiment with another P-gp over expressed cell line obtained from ATCC, USA (CRL-2274™). The current response increased with increase in the CRL-2274™ cells, however, the signals in this case was 17% lower compared to the CRL-1977™ cell type (data not
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Fig. 4. (A) Histogram showing the selective detection of MDRCC using biosensor at 30,000 MDRCC/mL concentration. (B) Microphotograph of MDRCC captured by ITO/ pTTBA/AntiP-gp probe ((a–c)) and (d) ITO/pTTBA chip treated with MDRCC (no cells were observed after washing the electrode).
shown). This is possibly due to the less concentration of the P-gp expressed on the cell membrane of CRL-2274™cells. 3.4. Selectivity, reproducibility, stability, and real sample analysis Control experiments were performed to validate the selectivity of the developed nanobiosensor. For this purpose, SKBr-3, HeLa, OSE, and HEK-293 cells were examined under the similar experimental conditions. The selectivity test was performed at 1000, 30,000, 75,000, and 100,000 cells/mL concentrations. Fig. 4 (A) shows the current response at 30,000 cells/mL concentration, where negligible signal is observed for tested cells. This was because no immunoreaction occurred between AntiP-gp and cell surface P-gp which was not present (or negligible) on these cells. We also performed control experiments to ensure the selectivity of the nanobiosensor by testing the compounds that are majorly present in the real sample matrix such as; albumin, fibrinogens, glucose. Interestingly, no current response was observed for these compounds indicating that the develop nanobiosensor is highly selective towards MDRCC detection and no interference or false
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positive signal is obtained. The reproducibility of the analysis was performed which showed the RSD o3.5% (n ¼5) and electrode-toelectrode RSD was o 3.2% even when the same preparation conditions were applied. This minor variation was possibly due to the difference in the modified surface and/or slight variations in the optimized conditions. We also studied the stability of the sensor probe with respect to time. The immunosensor retained almost 96% of its sensitivity for 6 weeks. The response decreased about 13% after 6 weeks and then gradually decreased with time. Thus, the sensor was stable for 6 weeks. The good stability was probably caused by the strong covalent attachment of the antibody on the stable conducting polymer–nanoparticle composite. The biomedical application of the nanobiosensor was investigated by detecting MDRCC in a human serum samples in the similar dynamic range (50–100,000 MDRCC/mL) as done in the buffer solution. Serum has been widely used as a model sample to demonstrate the biomedical application of biosensors (Chandra et al., 2013; Colin et al., 2011; Joshua et al., 2007). Thus, we used serum to examine the real clinical value of the developed sensor. MDRCC were dispersed into five times diluted serum samples and incubated with the GCE/AuNPs/pTTBA/AntiP-gp sensor probe. After washing the sensor probe was treated with the APBAMWCNT-Hyd conjugate, and reacted with H2O2 as described earlier. The current response increased with increase in the MDRCC number indicating that the developed biosensor can effectively detect MDRCC in serum samples, hence valuable in the biomedical analysis. The linear regression equation for the calibration plot in the serum sample is expressed as follows: ΔI ¼0.881 þ0.0023 [MDRCC] with the correlation coefficient of 0.998 indicating that this sensor is capable to detect MDRCC from the complex biological sample matrix effectively. The detection limit of MDRCC in serum sample was determined to be 2872 cells/mL (RSD o4.8%) based on the standard deviation of five repeated measurements of the blank (95% confidence level, n¼ 5). The detection limit in serum sample was 5 cells/mL higher (less sensitive) compared to the detection limit in a buffer solution (23 cells/mL). This minor variation in the detection limit was possibly due to negligible matrix effect due to the serum components. We also examined the applicability of the biosensor to detect MDRCC in the mixed cell sample. For this purpose, enumerated MDRCC were mixed with SKBr-3, HeLa, OSE, and HEK-293 cell and detected by the biosensor probe. Interestingly, the sensitivity of the detection in this case was 97% (n ¼5) compared when MDRCC was tested alone. This result clearly indicates that the developed biosensor can selectively detect MDRCC in the presence of other cancerous and noncancerous cells effectively. 3.5. Bioimagining and comparison with conventional reporter antibody based assay The capturing of the MDRCC by the AntiP-gp probe was investigated by bioimagining MDRCC incubated ITO/pTTBA/AntiP-gp surface using optical microscope. The ITO/pTTBA/AntiP-gp sensor was incubated with the MDRCC and after washing it was observed under microscope. Fig. 4(B) shows the microscopic images of ITO/ pTTBA/AntiP-gp sensor after treating it with different concentrations of MDRCC. With increase in the MDRCC concentrations between 102 to 103 cells/mL, the number of cells captured by the sensor was increased as shown in Fig. 4(B(a–c)). These results clearly indicate that ITO/pTTBA/AntiP-gp sensor is able to detect MDRCC. No MDRCC, however, were observed when Anti-gp was not immobilized on the ITO/pTTBA electrode indicating that AntiP-gp is essential to detect MDRCC (Fig. 4(B(d))). We also compared this method with the conventional reporter antibody based immunoassay. In this experiment, a reporter AntiP-gp-MWCNT-Hyd conjugate was prepared through the carbodiimide coupling
Fig. 5. Histogram showing the comparative response of the APBA-MWCNT-Hyd and reporter AntiP-gp-MWCNT-Hyd conjugate. The values on the y-axis shown in the histograms are based on the amperometric signals.
reaction following our earlier work (Zhu et al., 2010). This conjugate was tested by the following same experimental steps as described in the previous section. Fig. 5 shows the comparative response of the APBA-MWCNT-Hyd and the reporter AntiP-gpMWCNT-Hyd conjugate at 25,000, 50,000, and 100,000 MDRCC/mL. A much higher signal was observed for APBA-MWCNTHyd conjugate. The detection limit of MDRCC using reporter AntiPgp-MWCNT-Hyd conjugate was 15878 cells/mL based on the standard deviation of five repeated measurements of the blank (95% confidence level, n ¼5). This detection limit is 7 times higher i.e. less sensitive compared to the detection limit obtained using APBA-MWCNT-Hyd conjugate (237 2 cells/mL). This is due to the less immunoreaction between the reporter AntiP-gpMWCNT-Hyd conjugate and the MDRCC captured by the sensor probe. This less immunoreaction is due to the low concentrations of P-gp on the MDRCC cell membrane. It has already been reported that variation in the concentration of cell surface antigen (e.g.: P-gp in our case) significantly affects the antibody binding (Langmuir et al., 1991; Velders et al., 1998) which may directly influence the sensitivity of the biosensor. In case of APBA-MWCNTHyd conjugate low P-gp concentration does not affect the sensitivity of the biosensor because the target molecule for this conjugate is glycan which is expressed greatly onto the cancer cell surface (Zhang et al., 2010).
4. Conclusions A highly sensitive and selective amperometric nanobiosensor has been successfully designed for the detection of multidrug resistant cancer cells (MDRCC) in the biological matrix. The selectivity and sensitivity of MDRCC detection were achieved due to the monoclonal Pg-antibody covalently immobilized onto the conduction polymer-gold nanoparticle composite and also due the APBA-MWCNT-Hyd conjugate. The developed nanobiosensor is able to detect the MDRCC in mixed cell sample and in the presence of other chemical molecules. The detection limit of nanobiosensor is 23 72 MDRCC/mL, which is four times more sensitive than the most recent MDR cancer cell biosensor. The developed nanobiosensor is 7 times more sensitive compared to the conventional reporter antibody based immunosensor which gives a new direction towards sensitive detection of cancer cells. This method
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proposed herein is a generic method and can be easily adopted for other type of cancer cells simply by changing the detector antibody. The developed sensor could be a sensitive diagnostic method for the detection of drug resistant cancer cell in the biological fluids of cancer patients which will help clinicians to develop their alternate therapeutic strategies. The study clearly demonstrates highly sensitive detection of cancer cells even though the cell surface biomarkers is expressed/present in low concentrations using a nanobiosensor. In future, the developed nanobiosensor prototype can be translated into a miniaturized kit based assay for the point-of-care medical applications.
Acknowledgment This work was supported by the National Research Foundation of Korea (NRF) Grant funded by the Korea government (MSIP) (No. 20100029128).
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.2015.03.069.
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