Novel Graphene Biosensor Based on the Functionalization of Multifunctional Nano-bovine Serum Albumin for the Highly Sensitive Detection of Cancer Biomarkers
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A simple and convenient graphene bio-interface was designed by using multifunctional nano-denatured bovine serum albumin (nano-dBSA) film.
Highly sensitive cancer biomarker detection in diluted serum at the femtogram per milliliter level was achieved using the nano-dBSA functionalized graphene field-effect transistor.
KeywordsBio-interface Multifunctional denatured BSA GFET biosensor Cancer biomarker
Cancer is a major public health problem worldwide. For many cancers, it can take 20–30 years for initial lesions to progress to late-stage disease . Early detection is the key to cancer control, especially in reducing incidence rates and cancer-related deaths . Cancer protein biomarkers have been widely used in the early diagnosis of cancer. Carcinoembryonic antigen (CEA) is one of the most commonly used specific blood-based biomarkers for clinical tumor diagnosis. CEA is routinely used as an important indicator in annual medical checkups in many countries . Serum CEA concentration is closely correlated with malignant tumors, such as colorectal cancer , gastric cancer , medullary thyroid cancer , lung cancer , and pancreatic carcinoma . Determination of CEA concentration in a clinical sample can provide information about the severity of disease, tumor stage, pathological type, tumor metastasis, prognosis, and recurrence. Thus, CEA detection is valuable for the early diagnosis of cancer and has spurred efforts to develop strategies for the highly sensitive detection of CEA. The different strategies include photoelectrochemical immunosensors , time-resolved fluoroimmunoassay , surface-enhanced Raman scattering , fluorescence resonance energy transfer biosensors , electrochemical immunosensors , and electrochemiluminescence immunosensors [14, 15]. However, the development of a simple, low-cost, label-free, and rapid monitoring platform for the detection of cancer biomarkers for clinical diagnosis and screening applications remains a compelling goal.
Electrical detection of biomolecules based on their intrinsic charges is an efficient and ultrasensitive detection approach. Specifically, field-effect transistor (FET) biosensors are attractive because of their portability, inexpensive mass production, low power consumption, label-free detection, rapid response, and the potential for on-chip integration of the sensor and the electronic measurement system [16, 17]. In a FET biosensor, specific receptors immobilized in the sensing channel selectively capture the desired target biomolecules. The captured charged biomolecules can generate a doping or gating effect on the channel [18, 19, 20, 21]. Both are converted into a readable electrical signal by the FET, usually as a drain-to-source current or channel transconductance.
Interfacing biomolecules with channel sensing materials is a critical challenge to fabricate high-performance and inexpensive FET biosensors [22, 23]. In particular, the emergence of two-dimensional (2D) nanomaterials, such as graphene [24, 25, 26], molybdenum disulfide [27, 28], and black phosphorus , offers new powerful diagnostic tools for in vitro diagnosis and biomedical science applications. Graphene and graphene derivatives have been widely used in protein biomarker detection because of their tunable optical properties, high specific surface area, good biocompatibility, and easy functionalization [30, 31, 32, 33, 34]. Furthermore, the ambipolar field-effect, exceptional electrical properties, and atomically thin structures make graphene very promising as a channel material for FET biosensors , because of its excellent electrostatic coupling with charged target biomolecules.
The specificity and action of these biosensors depend on the coupling of effective recognition components on the graphene surface through noncovalent interactions that will not damage the graphene lattice or degrade its electronic performance. Noncovalent linkers mainly exploit π-stacking interactions and hydrophobic forces to attach directly on the graphene surface . Bifunctional noncovalent linkers, such as 1-pyrenebutanoic acid succinimidyl ester, N-hydroxysuccinimide (NHS) ester tripod, bovine serum albumin (BSA), pyrene butyric acid, and gold nanoparticles, have been successfully used to construct a bio-interface of graphene FET biosensors for the detection of glucose , DNA molecules , single-nucleotide polymorphisms , proteins [17, 39, 40, 41], and other biochemicals [42, 43]. Studies have focused on many difficult problems and topics. However, the complex and uncontrollable bio-interface of graphene FET channels remains a hurdle.
Herein, a simple and one-step method using multifunctional nano-denatured BSA (nano-dBSA) film to construct a graphene FET biosensor is described. The system enables the highly sensitive detection of cancer biomarkers. To construct the biosensor, native BSA protein solution was denatured by heating on graphene to form a layer that protected from unexpected destruction and surface contamination. At the same time, this nano-dBSA film could also serve as a cross-linker for the immobilization of anti-CEA monoclonal antibody (mAb). With the integration of the denaturation process into the fabrication of a graphene FET and the enriched chemical groups on the dBSA surface, one-step modification using 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC)/sulfo-NHS immobilized receptors on the graphene channel. In addition, enhanced sensitivity of the graphene FET biosensor was achieved by exploiting the dBSA modification method. Field-induced sensitivities to various CEA concentrations were observed, ultimately resulting in good specific recognition of CEA in diluted serum at an ultralow concentration of 337.58 fg mL−1. The cooperativity and strong affinity between CEA and anti-CEA mAb were estimated by the Hill model. The electric detection of the binding of CEA was interpreted to follow the Hill model for ligand–receptor interaction, indicating the negative cooperativity in binding between CEA and anti-CEA mAb with a dissociation constant of 6.82 × 10−10 M.
The demonstration of multifunctional nano-BSA chemical functionalization provides new functions for graphene-like 2D nanomaterials for further applications, such as biosensing, nanomedicine, imaging, cancer therapy, and drug delivery.
Graphene films grown by chemical vapor deposition on copper foil were purchased from 2D Carbon (Changzhou, China). BSA was obtained from Sangon Biotech (Shanghai, China, Purity: > 96%). The EDC and sulfo-NHS cross-linkers were purchased from Sigma-Aldrich (Darmstadt, Germany). Anti-CEA mAb1 and anti-CEA mAb2 were purchased from Medix Biochemica (Kauniainen, Finland). CEA protein and squamous cell carcinoma (SCC) were obtained from Fitzgerald (Acton, MA, USA) and Linc-Bio Science (Shanghai, China), respectively. Cytokeratin-19-fragment (CYFRA21-1) was purchased from Calbioreagents (Foster City, CA, USA). Quantum dots (QDs) with an emission wavelength of 625 nm were from Jiayuan Quantum Dot Co. (Wuhan, China). 1-Pyrenebutyric acid N-hydroxysuccinimide ester (PYR-NHS) was obtained from AnaSpec (Fremont, CA, USA). Polydimethylsiloxane (PDMS) was used to fabricate the reactive chamber. Deionized water obtained from a Millipore-Q purification system (Millipore, Billerica, MA, USA) was used for the preparation of all solutions.
2.2 Graphene Device Fabrication
Photoresist was used to define the drain/source electrode on a 300-nm SiO2/Si substrate, followed by the deposition of titanium and gold metals by electron beam evaporation. The metals on the photoresist were removed using acetone. The graphene-coated copper foil was etched using aqueous ammonium persulfate (10 g mL−1). Graphene was coated on the metal electrodes. Native BSA films were denatured on the graphene at 80 °C for 3 min, and the remaining dBSA/graphene films were etched using O2 plasma for 5 min. Finally, SU-8 photoresist was coated on the films as the insulating layer to prevent leakage current. The thickness of the dBSA functionalized graphene channel was optically characterized.
2.3 Bioprobe Functionalization and Characterization
The functionalization of dBSA on the graphene was carried out in a reactive chamber. The concentration of anti-CEA mAb used for immobilization onto the dBSA film was 2 mg mL−1. The dBSA film was incubated with 5 mg mL−1 EDC, 1 mg mL−1 NHS, and anti-CEA mAb solution in the dark. After incubation, the remaining unconjugated antibodies were removed by rinsing with phosphate buffered saline (PBS). A 1% BSA solution was used to block the excess reactive groups remaining on the graphene surface for 1 h. Secondary anti-CEA mAb was labeled with QDs (100 nM) mixed with 100 ng mL−1 CEA solution. The mixed solution was incubated with anti-CEA-dBSA functionalized graphene and bare dBSA functionalized graphene for 1 h each. Finally, the fluorescent images of each dBSA functionalized graphene were recorded using a fluorescence microscope.
All electrical measurements were performed using a semiconductor parameter analyzer (Keithley 4200).
3 Results and Discussion
3.1 Fabrication of Nano-BSA Graphene FET
3.2 Functionalization and Characterization of Bioprobes Based on Nano-dBSA Film
Sandwich fluorescent immunoassay is a commonly used approach in biotechnology . It was used to characterize the immobilization of anti-CEA mAb on dBSA film in this study. Secondary anti-CEA mAb conjugated with QDs was mixed with CEA solution and incubated with anti-CEA-dBSA functionalized graphene and bare dBSA functionalized graphene. Compared with the control group, the fluorescent images shown in Fig. S2 revealed that anti-CEA mAb was successfully immobilized on the dBSA functionalized graphene surface by the activation of EDC and sulfo-NHS. The results indicated that this novel method based on the dBSA film could be effective in the design of graphene biosensors.
3.3 Construction of Electrolyte-Gated Anti-CEA-dBSA Functionalized GFET
The reaction chamber made of polydimethylsiloxane was anchored on the substrate using silicone. The miniaturized Ag/AgCl electrochemical reference electrode was immersed in the reactive chamber as the gate of the anti-CEA-dBSA functionalized GFET. Drain–source voltage (Vds) and gate–source voltage (Vgs) were applied to force the operation of the devices. One terminal of the miniaturized Ag/AgCl electrochemical reference electrode was fixed on the shelf, and another terminal was immersed in the reactive chamber as the gate. Considering the sensitivity of anti-CEA-dBSA functionalized GFETs, 0.1 mM PBS was added to the reactive chamber as the electrolyte to maintain an appropriate Debye length . A schematic diagram of electrolyte-gated anti-CEA-dBSA functionalized GFET is shown in Fig. 2b. In addition, a representative optical micrograph of the dBSA functionalized graphene channel with an SU-8 insulating layer is shown in Fig. 2c.
3.4 Enhanced Performance of Anti-CEA-dBSA Functionalized GFET
The transconductance parameter gm for a transistor device is widely used to describe FET devices. This parameter represents the amplification capability of GFETs [26, 48], where a higher gm enables a greater conductivity response per unit of biomolecule charge excitation. Therefore, this parameter is positively correlated with the device sensitivity and is valuable for sensing applications. The transconductance gm of anti-CEA-dBSA functionalized GFETs under different drain–sources voltages is defined as the derivative of Ids with respect to Vgs in Fig. 3b. While the Vds was set at 0.1 V, gm = –577.78 μS approached the maximum (denoted gmax−) in the hole regime at a special gate voltage Vgs = 0.07 V (denoted Vmax-). Similarly, at Vgs = 0.24 V (denoted Vmax+), gm = 434 μS approached another maximum (denoted gmax+) in the electron regime. The average transconductance value of several anti-CEA-dBSA modified GFETs in Table S1 was higher than that of the anti-CEA mAb PYR-NHS modified GFETs in Table S2 and many other reported electrolyte-gated GFET devices [26, 49, 50], which revealed the high sensitivity of this device for biomolecule detection.
Using this model for the interfacial capacitance, the field-effect mobility of charge carriers in the device can be obtained. The mobility values extracted at the transconductance maximum points (gmax− for holes, gmax+ for electrons) were used as the mobility parameters of anti-CEA mAb modified GFET devices. Average values of hole mobility μave-h1 and electron mobility μave-e1 for seven anti-CEA-dBSA GFET devices were estimated to be approximately 2763.9 and 1169.6 cm2 V−1 s−1, respectively. As shown in Fig. 3c, d, the average mobility parameters of several anti-CEA mAb functionalized GFET devices based on a noncovalent functionalized linker (PYR-NHS) were lower than those of anti-CEA-dBSA GFET devices. These results indicated that GFET biosensors based on this multifunctional and self-protecting dBSA film could improve the performance of GFET biosensors.
3.5 Performance of Anti-CEA-dBSA Functionalized GFET
The drain–source current of the anti-CEA-dBSA functionalized GFET was monitored at various CEA protein concentrations to evaluate its sensing characteristics. The target CEA proteins at concentrations of 10 pg mL−1, 100 pg mL−1, 1 ng mL−1, 10 ng mL−1, and 100 ng mL−1 were introduced into the channel of the anti-CEA-dBSA functionalized GFET as the time-dependent response of the drain current was recorded (Fig. 4d). The mechanism of detection for anti-CEA-dBSA functionalized GFETs involved the adsorption of negative CEA proteins on the surface of the graphene. These proteins acted as electron donors, resulting in conductance changes. For this reason, the drain–source current increased gradually after injection of the target CEA at each concentration (Fig. 4d). According to the response of the control group, the limit of detection was less than 56 fM.
3.6 Target Detection in Diluted Serum Samples
According to the fitted red curve shown in Fig. 5b, ΔImax, Kd, and n were estimated to be 12.1 µA, 122.8 ng mL−1, and 0.35, respectively. The calculated value of n was less than 1, which indicated the negative cooperativity in binding interaction between CEA and anti-CEA mAb. The molecular weight of CEA of approximately 180 kD  resulted in a dissociation constant of 6.82 × 10−10 M. The dissociation constant between CEA and anti-CEA mAb had been investigated previously [58, 59], and it was determined to vary from 4 × 10−12 to 1 × 10−7 M. Therefore, the value of the resulting dissociation constant evaluated in this study using anti-CEA-dBSA functionalized GFETs was in accordance with previously reported results, indicating a high affinity between CEA and anti-CEA mAb. From Eq. (4) and the definition of the dissociation constant (Kd) , while the ligand concentration was equal to the dissociation constant (Kd), the percentage of bound receptors at equilibrium was 50%. According to the calculated value (122.8 ng mL−1) of Kd, the available receptors on the dBSA functionalized GFETs bio-interface were sufficient for the detection of CEA molecules under different concentrations in this study. According to the fitting results, the limit of detection was estimated to be approximately 337.58 fg mL−1, which was lower than for other graphene FET biosensors [41, 61, 62]. Well-defined drain–source current changes were observed for low CEA concentrations (337.58 fg mL−1) in diluted serum, which were much smaller than the cutoff value (5 ng mL−1) used in clinical diagnosis. In addition, compared with other nanomaterial-based CEA immunosensors in Table S3, the sensitivity of multifunctional dBSA functionalized GFETs showed obvious superiority. These results clearly demonstrated the promising potential of anti-CEA-dBSA functionalized GFETs in clinical applications.
A simple, convenient, and sensitive graphene–protein bioelectronic interface for GFETs based on a multifunctional nano-dBSA functionalized process was designed to target cancer biomarkers in diluted serum. This multifunctional nano-dBSA film formed on graphene acted as a protective layer and maintained the electronic properties of graphene during the fabrication of GFET devices and also served as a bifunctional cross-linker to bioconjugate anti-CEA mAb to detect CEA. This novel fabrication process made a high-performance GFET biosensor possible, as evidenced by electronic and fluorescent characterization. Good specificity and ultrahigh sensitivity (337.58 fg mL−1) toward CEA molecules were achieved by the measurement of drain–source currents of anti-CEA mAb functionalized GFETs. Measured responses with different orders of magnitude in analytes concentration displayed a good fit to a model based on the Hill binding equation, which indicated the negative cooperativity and a strong affinity between CEA and anti-CEA mAb binding interaction. Experimental results verified that the sensor response was derived from specific binding of the receptor to CEA, indicating that this multifunctional nano-dBSA film maintained its biologically active analyte-binding configuration when noncovalently bound to graphene. By functionalizing such different 2D nanomaterials with related receptors by this nano-dBSA process, it should be possible to offer controllable functionalization methods for various bio-interfaces for biosensors, nanomedicine, imaging, cancer therapy, and drug delivery.
The authors are grateful to the support of grants from the National Key R&D Program of China (Nos. 2018YFA0108202 and 2017YFA0205300), the National Natural Science Foundation of China (Nos. 61571429, 61801464, 61801465, and 81471748), the STS Project of the Chinese Academy of Sciences (NO.KFJ-STS-SCYD-120), the Science and Technology Commission of Shanghai Municipality (Nos. 16410711800 and 14391901900).
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