Specific biosensing using carbon nanotubes functionalized with gold nanoparticle–antibody conjugates

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Specific biosensing using carbon nanotubes functionalized with gold nanoparticle–antibody conjugates Shun Mao a, Ganhua Lu a, Kehan Yu a, Junhong Chen

a,b,*

a

Department of Mechanical Engineering, University of Wisconsin-Milwaukee, 3200 N Cramer Street, Milwaukee, WI 53211, USA College of Environmental Science and Engineering, State Key Laboratory of Pollution Control and Resource Reuse, The Institute for Advanced Materials and Nano Biomedicine, Tongji University, 1239 Siping Road, Shanghai 200092, PR China

b

A R T I C L E I N F O

A B S T R A C T

Article history:

This study demonstrates the electrical detection of protein binding by the introduction of

Received 18 July 2009

Au nanoparticle–antibody conjugates in a carbon nanotube field effect transistor (CNTFET),

Accepted 21 September 2009

in which the nanoparticle-functionalized carbon nanotube serves as the electrical conduct-

Available online 26 September 2009

ing channel. Antibody (anti-horseradish peroxidase) and antigen (horseradish peroxidase) binding events lead to the amplitude change in the drain current, which can be sensitively detected by FET measurements. The sensor shows negligible response to mismatched proteins such as Immunoglobulin G (IgG), confirming the specificity of the biosensor. The reported CNTFET-based biosensor could be adapted to detect a variety of proteins for in vitro diagnostics.  2009 Elsevier Ltd. All rights reserved.

1.

Introduction

Conventional biological sensing generally relies on optical [1,2], electrochemical [3], and piezoelectric [4] technologies, some of which can achieve detection limits from nanomolar to femtomolar levels [5,6]. Although these methods are sensitive and reliable for molecular detection in vitro diagnostics, they suffer from disadvantages such as inherent complexity and requirement for multiple reagents and steps, signal amplification, relatively large sample size, and high cost. Nanomaterials possess unique properties that are amenable to biosensor applications because they are extremely sensitive to electronic perturbations [7,8]. One-dimensional nanowires, especially silicon nanowires, have shown promising biosensing performance [9–11] based on their electrical sensitivity to the analyte-induced surface charges. By using smaller nanotubes with all atoms on the surface, biosensors based on carbon nanotubes (CNTs) have been demonstrated for biomolecular sensing [12–14]; and in particular, field effect transistors (FETs) based on semiconducting CNTs have been

used as biosensors [15–19]. These CNT-based FET devices are extremely sensitive to variations in the surrounding environment because all the electrical current flows through the outermost layer of the CNT which is in direct contact with the analyte. Several groups have reported on the electrochemical detection of protein binding [20,21] and DNA hybridization [22–25] using carbon nanotube field effect transistors (CNTFETs). The mechanism responsible for changing device characteristics (e.g., drain current amplitude and threshold voltage) is believed to be a charge-transfer reaction between analytes and CNTs; however, these approaches [21–23] require complex multiple-step and multiple-reagent procedures to label CNTs with probe proteins or DNAs. In addition, treating the CNT surface, especially the area around the CNT–metal contact, could modify device characteristics [26], which may confound the sensing results. Recently, the enhancement of the DNA detection by incorporating Au nanoparticles into CNTFETs has been reported [23,27]. In these studies, Au nanoparticles were linked with the target DNA to enhance

* Corresponding author: Address: Department of Mechanical Engineering, University of Wisconsin-Milwaukee, 3200 N Cramer Street, Milwaukee, WI 53211, USA. Fax: +1 414 229 6958. E-mail address: [email protected] (J. Chen). 0008-6223/$ - see front matter  2009 Elsevier Ltd. All rights reserved. doi:10.1016/j.carbon.2009.09.065

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the sensor performance; however, in addition to the probe and the target DNAs, a third DNA is needed to link the Au nanoparticle to the target DNA, which makes the fabrication and sensing processes more complex and difficult to control. Here we report on the fabrication and characterization of a biosensor based on CNTs decorated with Au nanoparticle– antibody conjugates. Au nanoparticles (Au NPs) labeled with anti-horseradish peroxidase (anti-HRP) were assembled onto the external surface of CNTs by a simple method that combines electrospray with electrostatic force directed assembly (ESFDA) [28–30]. Instead of being directly attached to the CNT, antibodies were engaged with CNT surfaces through the noncovalent binding of Au NPs to the CNT. Existing methods to assemble Au nanoparticles onto the CNT surface require the pretreatment of both the nanoparticle surface and the CNT surface through the coupling chemistry (e.g., polymers or linking proteins), which could introduce additional barrier for charge tunneling between the Au nanoparticle and the CNT and thus the optimum device performance may not be achieved. The elimination of unwanted proteins/chemicals due to the pretreatment also precludes side reactions that often complicate sensing signals. Using the same approach, the Au nanoparticle–antibody conjugates could be easily assembled onto various substrates, such as nanotubes, nanowires, and graphene sheets, which implies that the reported sensor design and fabrication method is potentially useful for many different types of biosensors. In addition, the distribution and the concentration of probe antibody can be controlled by simply adjusting the areal density of Au NPs on CNTs, which modulates the upper detection limit for sensing analytes at various concentration levels. The CNTFET-based sensor was capable of recognizing the target antigen, HRP, through Au NPs labeled antibody (anti-HRP) with a high degree of control and specificity. Antibody–antigen can be detected by measuring the change in the CNTFET characteristics, from which the detection limit of the sensor was determined to be on the order of lM. Immunoglobulin G (IgG) was used as a mismatched protein to verify the specificity of the sensor. Bovine serum albumin (BSA) was used to block the nonspecific binding of proteins with the sensing element.

layer of thermally-formed SiO2 (thickness of 200 nm). Multiwalled CNT (MWCNT) powders were first dispersed in the DI water (0.005 mg/ml) with aqueous dispersant from Alfa Aesar. After ultrasonication for 30 min, one drop of the MWCNT suspension was pipetted onto the electrodes, then dried under room temperature. The device was then annealed in argon flow (1 lpm) for one hour at 200 C to remove residue solvents and improve the contact between CNTs and electrodes. After annealing, it was found that the device was very robust. Carbon nanotubes were immobilized between two metal electrodes and could not be washed away after several cycles of washing and drying, which was confirmed by the scanning electron microscopy (SEM) imaging. Au nanoparticles labeled with anti-HRP or bare Au NPs were assembled onto the surface of CNTs through a simple method that combines electrospray with an electrostatic force directed assembly (ESFDA) technique. The assembly time was one hour for Au NP–anti HRP conjugates and Au NPs. Before introducing HRP, CNTFETs were modified with BSA (0.01 mg/ml) to reduce the possible nonspecific binding of HRP to CNTs and electrodes. Devices were incubated with BSA for 2 h at room temperature and then washed with the PBS buffer. After that, 0.02 ml HRP was pipetted onto the device for protein binding for one hour, followed by washing and drying. FET measurements were carried out using a Keithley 2602 source meter by recording the drain current response (Vd = 0.01 V) when ramping the gate voltage Vg from 3.0 to 3.0 V (with a step of 0.01 V). FET measurements were performed at various steps during the device fabrication and test. The sensor reproducibility/repeatability was studied by using 3–4 sensors in each group of sensing tests. It was found that the sensors had similar sensing responses for the same test, which indicated that our sensor responses were reproducible. For our sensor structure, it is difficult to remove the HRP from the Au conjugates after the sensing process. Therefore, the same sensor cannot be recovered/reused for multicycle sensing. A Hitachi S-4800 UHR FE-SEM was used for SEM characterization of the biosensing device at an acceleration voltage of 30 kV.

2.

Fig. 1 shows a schematic diagram of the CNTFET, in which anti-HRP were anchored to the CNT surface by Au NPs and functioned as specific recognition groups for HRP binding. The CNTFETs were fabricated by creating a random network of CNTs between Au electrodes through bottom contact geometry. In this configuration, Au electrodes were used as metal contacts for drain and source; the SiO2 layer was used as the insulating material, through which a gate bias was applied to the back silicon wafer during the FET measurement. This nanoscale FET device, with CNTs as the conducting channels, could be used to detect biomolecule binding [16,21]. The sensing response was attributed to the change in the electric characteristic of the conducting channels, i.e., CNTs, presumably due to the effective electronic transfer between the Au nanoparticle and the CNT. In this paper, we examine the dependence of the drain current Id on the gate voltage Vg, referring to this response as the ‘‘device characteristic,’’

Experimental

Au nanoparticles labeled with anti-HRP (18 nm colloidal gold AffiniPure Goat anti-HRP) were purchased from Jackson Immuno Research and used without further purification. Bare Au NPs (20 nm colloidal gold) were purchased from BB International. HRP and IgG from human serum were ordered from Sigma–Aldrich. The blocking agent BSA was purchased from Rockland. PBS (pH 7.4, ·1) (Fisher BioReagents) was used as the solvent for HRP, IgG, and BSA. All solutions were prepared with distilled and deionized water (DI water) supplied by Cellgro. Multi-walled CNTs with diameter of 20 nm and lengths of several microns were purchased from Alfa Aesar. Gold interdigitated electrodes [31] with both finger width and inter-finger spacing (source–drain separation) of about 1 lm were fabricated using an e-beam lithography process (Raith 150 lithography tool, 30 kV) on an Si wafer with a top

3.

Results and discussion

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Fig. 1 – Diagram of a CNTFET. Anti-HRP is anchored to the CNT surface through Au NPs and functions as a specific recognition group for HRP binding.

as a result of various protein binding events. The devices displayed p-type semiconducting behavior in an ambient environment, presumably due to the exposure of CNTs to oxygen (O2 + 2e = 2O ), as other groups also have observed [16,17,32]. The dependence of the drain current Id as a function of the drain–source bias Vd has also been recorded to understand the sensor response and to verify the efficacy of the BSA blocking. Procedures to fabricate the CNTFET-based biosensor are schematically shown in Fig. 2a. A network of CNTs was first obtained by dispersing CNTs onto interdigitated electrodes, followed by the antibody attachment through the noncovalent attachment of antibody-conjugated Au NPs to CNT surfaces. A schematic diagram of the assembly of Au NP–anti HRP conjugates is shown in Fig. 2a (dashed frame). Colloidal Au NPs with diameters of 18 nm and anti-HRPs on their surfaces were aerosolized by an electrospray process and assembled onto the surface of CNTs through ESFDA [28–30]. This procedure has been successfully used for decorating CNTs with various nanoparticles [30,33] and we found that the antibody activity was preserved after the electrospray assembly process. After Au NP–antibody conjugates were assembled onto the CNTs, the sensing device was modified with BSA by incubating the device with BSA for two hours at room temperature. BSA is expected to cover all the surface area of the device except the antigen binding sites; therefore, it can reduce the possible nonspecific binding of the HRP to CNTs and electrodes. In the final step, HRPs were pipetted onto the device for protein binding experiments. The electrical characteristics of the sensing device were measured and compared before and after each of the above-mentioned procedures. To confirm the presence of Au NPs on the surface of the CNT, SEM was used to examine the CNTFET before and after the electrospray of Au NP–antibody conjugates. Fig. 2b shows a single bare CNT with a diameter about 20 nm and length of several lm spanning between two Au electrodes, working as the conducting channel for the FET. After being electrosprayed with Au NP–antibody conjugates, many Au NPs (light dots) were uniformly distributed on the surface of the CNT (Fig. 2c), indicating the addition of Au NPs into the CNTFET. Based on the SEM images, we conclude that Au NPs have been effectively attached to the CNTs likely through noncovalent

binding [28,30]. Since anti-HRPs are covalently linked to the surface of Au NPs, we assume that the antibodies are present in the sensor for probing antigens. This assumption has been confirmed by the sensing results presented below. Typical gate voltage dependence of the normalized drain current Id (Id is normalized by the drain current of the device at Vg = 3.0 V) after CNTFETs are treated with Au NP–anti-HRP conjugates and BSA is shown in Fig. 3a; further treatment with PBS buffer and HRP is shown in Fig. 3b. During the FET measurement, a gate bias Vg ramped from 3.0 V to 3.0 V while a constant bias of 0.01 V was held between the drain and the source. From Fig. 3a, the gate voltage dependence of the normalized drain current Id for the bare CNT is typically p-type, in which holes are the majority carriers and the conductivity of the CNT depends on the mobility and density of holes. When a negative bias (Vg) is applied to the Si wafer, an electric field is built across the insulation layer (SiO2 layer); leading to the hole accumulation on the top layer of SiO2 and the subsequent hole transfer to the CNT through Au electrodes. In this case, the density of holes in the CNT increases, leading to the raise of Id. On the other hand, a positive Vg results in the electron transfer to the CNT and the decrease in Id. When the gate bias reaches the threshold point, the FET is off and Id decreases to the smallest value; then further increase of the Vg will change the device type from p-type to n-type and Id will increase accordingly. From Fig. 3a, the type of the device did not change after the CNTs were coated with Au NP–anti-HRP conjugates; however, the drain current Id decreased. There are two competing factors that could contribute to the observed phenomenon. Firstly, after being attached onto the CNTs, Au conjugates increase the scattering centers across the tube; thereby reducing the mobility of holes in the CNT and thus leading to the decrease in the drain current. Secondly, by comparing the work function of Au NPs (5.1– 5.47 eV) [34] and CNTs (4.3–4.9 eV) [35–37], electrons may transfer from the CNT to the Au NP, increasing the density of holes in the CNT and thereby increasing the drain current. Based on the decrease of the drain current, we suggest that the decrease of hole mobility in the CNT is the dominant factor. The drain current Id further decreased after the CNTFETs were treated with BSA, possibly because the BSA was negatively charged when dissolved in a phosphate buffered saline

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Fig. 2 – (a) Schematic illustration of the fabrication process of the CNTFET-based biosensor with the assembly of Au NP–anti HRP conjugates onto the CNTFET shown in the dashed frame. (b) SEM image of a single bare CNT between electrodes. (c) SEM image of a single CNT coated with Au NP–anti HRP conjugates and spanning across an electrode gap.

(PBS) buffer and negative charges could transfer from the BSA to CNTs. Fig. 3b demonstrates a significant sensing response as a result of antigen–antibody binding, which is evidenced by a decrease of Id by about 40% (at Vg = 3.0 V) with the presence of HRP. A control experiment was performed by introducing only the PBS buffer to the device and a very minute response (5% decrease in Id) from the sensor was observed, indicating that the sensor response is mainly due to the binding of HRP to anti-HRP. The analyte–CNT interaction may have two effects on the CNT conductance [16]. Firstly, the analyte (HRP) triggers various scattering centers across the CNT as a result of binding events. Upon anti-HRP–HRP binding, geometric deformations occur and increase the scattering centers across the tube; thereby reducing the mobility of holes in the CNT and thus leading to the decrease in the drain current. As a possible second effect, the decrease of the drain current could also be attributed to the charge tunneling from the anti-HRP to

CNTs as a result of binding events. HRP is a glycoprotein with four lysine residue, in which lysine is an a-amino acid with the chemical formula HO2CCH(NH2)(CH2)4NH2. Charge tunneling through HRP to anti-HRP is less likely since both HRP and anti-HRP are insulator and the antibody–antigen binding is a noncovalent interaction; however, charge tunneling from a-amino acid groups in HRP to the CNT through the solvent is possible in the PBS buffer solution (conductivity 14–16 mS/ cm), which leads to a decreased charge carrier (hole) density in CNTs and thus decreased drain current. The observed conductivity change in the Au/CNT structure could be attributed to both the scattering and the charge tunneling mechanisms [21,23]. Further investigation is needed to identify the dominant mechanism. Two control experiments were performed to better understand the sensor response to HRP and verify the blocking efficacy of BSA, respectively. In the first control experiment, we performed the sensor test with bare Au NPs instead of the

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Fig. 3 – Typical gate voltage dependence of the normalized drain current Id for CNTFETs treated with (a) Au NP–anti-HRP conjugates and BSA; (b) BSA, PBS buffer, and HRP. (Vd = 0.01 V, HRP concentration = 4.5 lM.) Drain–source voltage dependence of drain current Id for CNTFETs treated with (c) Au NPs, BSA, and HRP; (d) Au NPs and HRP.

Au NP–anti-HRP conjugates. Drain–source voltage dependence of drain current Id is given in Fig. 3c. It is well known that the ideal metal–semiconductor contacts are Ohmic, which have a linear I–V characteristic in both biasing directions. From the I–V characteristics in Fig. 3c, it was found that drain–source voltage dependence of Id after being treated with bare Au NPs were slightly non-linear, indicating that there was a Schottky barrier present at the metal/CNT interface [26]. To reduce the influence of the Schottky contact on the sensing response, we performed an annealing treatment to the device in argon flow and found that Schottky contacts could transform to Ohmic contacts after annealing. From Fig. 3c, we found that the drain current increased after bare Au NPs were assembled onto the surface of the CNTs, which can be explained by the electron transfer from CNTs to Au NPs, as previously discussed. After the BSA blocking, the drain current decreased as expected; however, in this case, there was no change in the drain current after the HRP was introduced to the sensor. As we designed, if there was no antibody labeled on the Au NPs, there were no binding sites for the antigen; therefore, there was no response from the device. This proves that linking antibodies to the CNT through Au NPs is a key step for the function of the sensor. The result also confirms that the sensing response in the device (Fig. 3b) comes from the binding of HRP to anti-HRP. To verify the blocking efficacy, no BSA treatment was applied to the device before introducing HRP in the second

control experiment. The drain–source voltage dependence of drain current Id is given in Fig. 3d. From the drain current measurement, we found that the drain current decreased by about 40% at Vd = 2.0 V when HRP was added to the device, which came from the nonspecific binding of HRPs to CNTs. In this case, with no BSA on the device, HRP could attach to the whole surface of CNTs, which increased the scattering centers across the tube and decreased the conductance of the device. The results in Figs. 3c and d suggest that BSA can effectively block the nonspecific binding of HRP to the sensing element. Thus, treating the device with BSA is necessary for the specific function of the sensor and it can effectively diminish the side response from the nonspecific binding of the analyte to the device. Very often, sensors are useful only when there is specificity. To verify the specificity of the sensor, Immunoglobulin G (IgG) was used as a mismatched protein. Fig. 4a summarizes the relative increase in the sensor resistance (sensitivity) upon exposure to the target protein HRP, the mismatched protein IgG, and the PBS buffer. The increase in the sensor resistance due to the mismatched IgG (5.5%) or PBS buffer (2.7%) is significantly smaller than that caused by the HRP (32.7%). Fig. 4b shows the relationship between the HRP concentration and the sensor sensitivity, which increases with the increase of the HRP concentration. Fig. 4b also indicates that a 0.45 lM complementation HRP detection (7.6% resistance increase) can be sufficiently differentiated from the mismatched IgG and the PBS buffer.

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Fig. 4 – (a) Comparison of the CNTFET sensitivity (relative change in the sensor resistance) induced by anti-HRP binding with target HRP (4.5 lM), with mismatched IgG (4.5 lM), and PBS buffer. (b) The biosensor sensitivity as a function of the HRP concentration. Error bars were obtained through the error propagation.

We conclude that the detection limit of the sensor is on the order of lM based on the biosensor sensitivity summarized in Fig. 4b and this could be improved further by optimizing CNTs, Au NPs, and the metal contacts. The CNT concentration, the arrangement of CNTs in the FETs, and the size and the areal density of Au NPs can influence the sensor performance. For example, single or fewer CNTs have been shown to enhance the sensitivity and detection limit of the sensor. Dong et al. reported that devices made of Ta electrodes could have a more pronounced response than those with Au electrodes [23], indicating that the Fermi energy difference between the electrode and the CNT may affect the interface change upon the exposure of analyte to the device. Also, it was reported [16,22] that the metal/nanotube interface or contact region is highly susceptible to modulation by adsorbed species. The modulation of the metal work function can alter the Schottky barrier of the metal/nanotube interface, thereby leading to a significant change in the nature of contacts and consequently a change in the conductance of the device. To fully understand the sensing mechanism, the

contribution from analyte adsorption to the metal/nanotube contact region will be investigated in future work.

4.

Conclusions

In summary, we have successfully demonstrated an electrical detection of protein binding in CNTFETs by the introduction of Au NP–antibody conjugates to the surface of CNTs. The protein binding events can be sensitively detected by the amplitude change in the drain current from FET measurements, through which proteins with a concentration on the order of lM can be accurately detected. The device sensitivity can be further improved by using single-walled CNTs (SWCNTs), through adjusting the composition of the CNTs (ratio of semiconducting to metallic tubes), or by using various electrode materials. In addition, this sensor can be used for sensing different analytes by decorating CNTs with different Au NP–antibody conjugates to realize the specificity, which makes it capable of detecting a variety of proteins for in vitro diagnostics.

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Acknowledgements Financial support for this work was provided by the National Science Foundation (CMMI-0900509, CBET-0803142 and ECCS0708998). The SEM imaging was conducted at the Electron Microscope Laboratory of UWM. The e-beam lithography was performed at the Center for Nanoscale Materials of Argonne National Laboratory, which is supported by the US Department of Energy, Office of Science, Office of Basic Energy Sciences, under Contract No. DE-AC02-06CH11357. The authors thank Prof. D.P. Klemer for insightful discussion and helpful comments, Dr. H.A. Owen for technical support with SEM analyses, and Dr. L.E. Ocola for assistance in the electrode fabrication. The authors also thank anonymous reviewers for valuable comments and suggestions.

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