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DNA sensors based on CNT-FET with floating electrodes
Sensors and Actuators B 169 (2012) 182–187
Contents lists available at SciVerse ScienceDirect
Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb
DNA sensors based on CNT-FET with floating electrodes Byeongju Kim a , Joohyung Lee a , Seon Namgung a , Jeongsu Kim a , Jae Yeol Park b , Moon-Sook Lee c , Seunghun Hong a,d,∗ a
Department of Physics and Astronomy, Seoul National University, Seoul 151-747, Republic of Korea Department of Automotive Engineering, Doowon Technical University College, Anseong 456-718, Republic of Korea Bio Center, Samsung Advanced Institute of Technology, Yong-In 446-712, Republic of Korea d Department of Biophysics and Chemical Biology (WCU Program), Seoul National University, Seoul 151-747, Republic of Korea b c
a r t i c l e
i n f o
Article history: Received 24 December 2011 Received in revised form 14 April 2012 Accepted 22 April 2012 Available online 28 April 2012 Keywords: Carbon nanotube Sensor Floating electrode Schottky barrier DNA
a b s t r a c t We report a successful development of a floating electrode-based DNA sensor with controllable responses. Here, metallic floating electrodes were fabricated to form Schottky barriers between carbon nanotubes and the floating electrodes. We showed that the sensor response could be enhanced by increasing the number of floating electrodes. We also analyzed the response of the sensors based on the Langmuir isotherm theory. © 2012 Elsevier B.V. All rights reserved.
1. Introduction The electrical detection method of DNA can have several advantages over bulky fluorescence-based DNA detection method, such as label-free detection, high sensitivity, and easy integration with other devices for practical applications [1–6]. Recently, diverse DNA sensors based on nanomaterials have been developed. Among them, carbon nanotubes (CNT) network-based DNA sensors have attracted much attention due to its high sensitivity. In the CNT network-based DNA sensors, the modulation of Schottky barriers by DNA hybridization has been known to play a crucial role in detecting target DNA. When the target DNA was bound to probe DNA, the height of Schottky barriers at the contact regions between CNTs and metal electrodes were varied. Since the height of Schottky barriers controls the electric currents of the devices, the binding of target DNA could be detected by monitoring the change of the electric currents in the devices. There has been much effort to enhance the sensitivity of the Schottky barrier-based sensors by forming a nonsymmetrical Schottky contact [7], by increasing Schottky contact area using a shadow mask [8], and by varying the contact metal
∗ Corresponding author at: Department of Physics and Astronomy, Seoul National University, Seoul 151-747, Republic of Korea. Tel.: +82 2 880 1343; fax: +82 2 884 3002. E-mail address: [email protected] (S. Hong). 0925-4005/$ – see front matter © 2012 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.snb.2012.04.063
[9]. However, in the previous methods, it was difficult to predict or to control the sensor responses for practical applications. Herein, we developed a floating electrode-based sensor for the detection of DNA molecules with controllable responses. We showed that sensor responses could be enhanced simply by increasing the number of floating electrodes. The experimental results were explained by the modulation of Schottky-barriers formed at the floating electrodes. We also demonstrated that experimental results fitted well to the Langmuir isotherm theory. Our approach could be a simple but efficient way to improve the sensitivity of CNT network-based sensors and should provide an important insight regarding Schottky barrier-based sensors in general.
2. Experimental 2.1. Materials Purified single-walled carbon nanotubes (swCNTs) were purchased from Carbon Nanotechnologies Inc. (USA) and used as received. The swCNTs have a diameter of 0.7–2 nm and a length of 2–3 m. Mercaptohexanol (MCH), octadecyltrichlorosilane (OTS) and other chemical reagents were purchased from Sigma–Aldrich (USA) and used as received. DNAs used in these experiments were purchased from GenoTech Company (Korea). Their base sequences are listed below:
B. Kim et al. / Sensors and Actuators B 169 (2012) 182–187
16 mer probe DNA: Thiol-C6 -5 -GCC ACA AAC ACC ACA A-3 16 mer target DNA: 5 -TTG TGG TGT TTG TGG C-3 16 mer noncomplementary DNA: 5 -CAT GGT TGA TCC GTT C-3 25 mer probe DNA: Thiol-C6 -5 -AGA CCT CCA GTC TCC ATG TTA CGT C-3 25 mer target DNA: 5 -GAC GTA ACA TGG AGA CTG GAG GTC T-3 25 mer noncomplementary DNA: 5 -ACG CTG AGT ACG GGT GCA AGA GTC A-3 2.2. Fabrication of CNT-field effect transistor (CNT-FET) devices To create CNT channel patterns, we first patterned AZ 5214 photoresist (PR) on a silicon oxide wafer via standard photolithography. Then, the wafer was placed in an OTS solution (1:500, v/v in hexane) for 5 min. Subsequently, the PR patterns were removed using acetone. Then, the patterned silicon oxide wafer was placed in a CNT solution (concentration 0.1 mg/ml) for 30 s, which was prepared by dispersing CNTs in 1,2-dichlorobenzene with ultrasonication for 1 h. This step allowed CNTs to be adsorbed selectively onto bare SiO2 regions in the wafer, while the methyl-terminated OTS selfassembled monolayer (SAM) blocked non-specific adsorption of the CNTs. The wafer was rinsed thoroughly with 1,2-dichlorobenzene and dried with nitrogen gas. After then, metal electrodes (30 nm Au on 10 nm Ti) were formed via standard photolithography, thermal evaporation, and lift-off process. Finally, we covered the source and drain electrodes with a photoresist layer to eliminate leakage currents from the electrodes in a buffer solution. 2.3. Functionalization of electrodes with probe DNA First, the CNT-FETs were placed in a probe DNA solution (10 M in phosphate buffered saline (PBS, pH 7.4, Fisher Scientific)) for 10 h to functionalize the Au electrodes with the probe DNA. The devices were rinsed with PBS and dried with a N2 stream. The devices were placed in MCH solution (10 mM in deionized water) overnight and dried with a N2 stream. 2.4. Electrical detection of target DNA using CNT-FETs In DNA detection experiments, the source–drain currents of the floating electrode-based CNT-FET devices were measured by a semiconductor characterization system (Keithley, 4200, USA). The system was connected to a probe station that makes electrical contacts to the source and drain electrodes of the CNT-FET devices. During the measurement process, the source–drain (SD) bias (Vsd ) of 100 mV was applied. A PBS buffer solution was dropped onto the channel regions of the devices. Then, an Ag/AgCl electrode, used as a gate electrode, was inserted into the solution, and a gate voltage (Vlg ) of 0 V was applied. The Ag/AgCl electrode was used to minimize the side effects such as electrochemical reactions at the reference electrode–solution interface [10]. After the base level of the signal was stabilized, the liquid gate characteristics of the device were measured. For liquid gating measurements before DNA hybridization, the gate voltage was swept from −0.5 to +0.5 V with a Vsd of 100 mV. Then, for the detection of DNA hybridization, target DNA solutions with various DNA concentrations, from 100 fM to 10 M, were introduced into the buffer solution. 3. Results and discussion 3.1. Sensing mechanism Fig. 1a illustrates the structure of our floating electrode-based DNA sensor (upside) and its corresponding energy-band diagram (downside). Here, probe DNA was adsorbed only onto the surface of the floating electrodes. Source and drain electrodes were covered
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by photoresist to block leakage currents. The energy-band diagram shows the formation of Schottky barriers at the contact regions between CNTs and metal electrodes. These Schottky barriers determine the conductance of the device [11,12]. Fig. 1b shows the schematic diagram of the target DNA sensing experiment. When a target DNA solution was injected, the target DNA molecules were hybridized with their complementary probe DNA molecules on the floating electrodes (upside). During the hybridization process, the work function of the floating electrodes was decreased, and the Schottky barrier height was increased (downside). Since this increased Schottky barrier was acting as a potential barrier for hole-carriers, the currents in the channel region were decreased [13,14]. Note that the electrical conductance of device was determined solely by the floating electrodes because source and drain electrodes were passivated with PR. Therefore, the capability of the devices to detect target DNA could be controlled simply by varying the number and the shape of the floating electrodes. 3.2. Device layout and electrical characterization of CNT-FETs Fig. 2a shows the channel region of our floating electrode-based CNT-FET devices. The scanning electron microscopy (SEM) image (left) clearly shows that each floating electrode has dimensions of 12 m width and 2 m length. Here, the CNT network channel has dimensions of 2 m width and 40 m length. Note that the size and the number of floating electrodes can be readily controlled by simple photolithography processes. The right image of Fig. 2a is an atomic force microscope (AFM) image showing the CNT channel region between floating electrodes. Note that the CNTs were assembled only on a bare SiO2 region. The electrical properties of the floating electrode-based CNTFETs were investigated. Fig. 2b shows the typical liquid gating effect of our floating electrode-based CNT-FETs. The liquid gating effect was measured by applying a liquid gate bias (Vlg ) to the PBS buffer solution on a floating electrode-based CNT-FET. Here, an Ag/AgCl electrode was used as a liquid gate electrode, and a gate bias was swept from −0.5 V to 0.5 V with a source–drain bias of 0.1 V. The graph shows that the device also had p-type characteristics in the PBS buffer solution. Since the electric currents were drastically changed in the small range of voltage sweep, this device could be utilized as a sensitive sensor. We also analyzed the noise characteristics of our devices. The inset of Fig. 2c shows a typical noise spectrum of our devices with various numbers of floating electrodes at low frequency ranges. The graph shows a typical 1/f noise behavior, and the noise spectrum can be fitted by SI ∼ 1/fˇ with ˇ values of 1. The graph also shows that the noise amplitude was increased as the number of floating electrodes was increased. We confirmed that the noise amplitude A is linearly proportional to the device resistance R following the relationship of A ≈ 10−11 × R, which is a typical behavior of CNT networks (Fig. 2c) [15–17]. This result indicates that the noise of the floating electrode-based devices was dominated by CNT-networks in the channel region. 3.3. Detection of target DNA using floating electrode-based CNT-FETs Fig. 3a shows the response of our DNA sensors to the injection of non-complementary and complementary DNAs. When non-complementary DNA was injected, there were no significant changes in electric currents. However, when complementary DNA was injected, the significant decrease of electric currents was observed. The reduction of the electric currents can be explained by the hybridization of complementary target DNA molecules with the probe DNA molecules fixed on floating electrodes. The hybridization resulted in the increase of the height of Schottky barrier, which consequently reduced the currents of the devices (Fig. 1) [12,18].
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Fig. 1. Sensing mechanism of floating electrode-based CNT-FET sensors. (a) (Upside) Schematic diagram of a fabricated device before target DNA binding. Probe DNA was adhered onto a floating electrode surface, and source and drain electrodes were covered with photoresist to block leakage currents. (Downside) Energy-band diagram of the floating electrode-based CNT-FET device. Schottky barriers were formed at contact regions between CNTs and metal electrodes. Here, hole-carriers passed through the Schottky barriers formed at the floating electrode. (b) (Upside) Schematic diagram of target DNA sensing. The target DNA was hybridized with probe DNA on the surface of the floating electrode. (Downside) Energy-band diagram of the floating electrode-based CNT-FET device after target DNA binding. When the target DNA was hybridized with probe DNA, the work function of the floating electrode was decreased, which increased the height of Schottky barriers. Due to the increased height of Schottky barriers, the currents of the device, dominated by hole-carriers, were decreased.
Fig. 2. Characterization of floating electrode-based CNT-FETs. (a) SEM (left) and AFM (right) image of a floating electrode-based CNT-FET which has seven floating electrodes. Individual floating electrode had dimensions of 12 m width and 2 m length, while a CNT network channel had dimensions of 2 m width and 40 m length. (b) Liquid gating effect of a floating electrode-based CNT-FET. Here, an Ag/AgCl electrode was used as a liquid gate, and the gate bias was swept from −0.5 V to 0.5 V with a source–drain bias of 0.1 V. (c) Graph showing the noise amplitude (A) versus the channel resistance (R) relation for the floating electrode-based CNT-FETs with various numbers of floating electrodes. A fitting line represents A ≈ 10−11 × R. The insert shows the noise power spectrum (SI ) of the floating electrode-based CNT-FETs with various numbers of floating electrodes at low frequency ranges.
B. Kim et al. / Sensors and Actuators B 169 (2012) 182–187
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Fig. 3. Response of a floating electrode-based CNT-FET sensor to the binding of target DNA to probe DNA. (a) Real-time response of a floating electrode-based DNA sensor to the injection of DNA. A drastic change of source–drain currents was observed only when complementary target DNA was introduced. Arrows indicate the points of DNA injections. (b) Real-time response of a floating electrode-based DNA sensor to the injection of target DNA with various concentrations. The significant reduction of currents was observed on the addition of target DNA from 1 nM to 10 M concentration. Arrows indicate the points of target DNA injections. (c) Response curves of floating electrodebased DNA sensors with various numbers of floating electrodes. Here, the sensitivity of floating electrode-based DNA sensors was defined as G/G0 . Note that the sensitivity of the sensors at certain concentration was increased as the number of floating electrodes was increased. (d) Response curves of floating electrode-based DNA sensors to 16-mer and 25-mer target DNAs. Note that the detection limit for 25-mer DNA was 100 fM, while that for 16-mer DNA was 100 pM.
This result clearly shows that our floating electrode-based CNTFETs could selectively detect the complementary target DNA in real time. When we functionalize floating electrodes by probe DNA molecules, the probe DNA molecules could wrap CNTs. These DNA molecules are known to play little role in hybridization with complementary DNA, while they block the non-specific adsorption of analytes [14]. Fig. 3b shows the real time measurement of the source–drain currents in our sensor device during the introduction of target DNA molecules with various concentrations ranging from 1 nM to 10 M. The addition of target DNA molecules caused a sharp decrease in the source–drain currents, and then the source–drain currents were gradually saturated. The change of currents (G/G0 ) increased as the concentration of target DNA solutions was elevated. The sensor response could be analyzed by a model based on the Langmuir isotherm theory [19]. If we assume that the maximum surface density of target DNA on the floating electrodes was Cs max and that target DNA molecules were hybridized with probe DNA molecules following the Langmuir isotherm, the surface density Cs of target DNA molecules bound to probe DNA molecules can be written as Cs =
Cs max · C 1/K + C
where C and K are the concentration of target DNA molecules in the solution and the equilibrium constant between probe DNA and target DNA, respectively. Furthermore, if we assume that the sensor signal (G/G0 ) can be approximated as |G/G0 | ∼ kCs where k is
a constant representing the characteristics of sensor device, the sensitivity of the sensor can be written as
Cs max · C G G = k 1/K + C 0
where G0 and G represent the original conductance of device and the conductance change of the sensor, respectively. From the above equation, we could deduce that the both parameters of k and K have great importance related to the sensitivity of the sensor device where k and K are related to the number of floating electrodes and the binding between target DNA and probe DNA, respectively. Fig. 3c shows the response of the sensor devices with various numbers of floating electrodes. Here, the sensitivity of floating electrode-based DNA sensors was defined as G/G0 . Using the above theoretical model, we performed curve fitting to experimental data sets with a large concentration range. The graph clearly shows that the experimental results fit well with the theoretical model. In previous works, we also applied the similar equation for the estimation of the equilibrium constant for the binding of mercury ions onto a CNT surface [19]. Here, all the equilibrium constants of probe DNA to target DNA with various numbers of floating electrodes had values in the same order (∼107 M−1 ). This result indicates that the equilibrium constant for given pair of DNA molecules did not change even though the structure of the sensors was varied. Note that the sensitivity of the sensors at a certain concentration was increased as the number of floating electrodes was increased. For example, the sensitivity of the sensor at 1 nM concentration with one floating electrode was 8%, while that with seven floating electrodes was 21%. As mentioned earlier, the
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sensor response was controllable by varying the number of floating electrodes. We performed control experiments to check if the enhanced sensitivity of our sensors with larger number of floating electrodes came simply from the increased surface area of the floating electrodes rather than the increased number of the floating electrodes (Fig. S1 in Supporting Information). Here, we compared the sensitivity of a sensor including seven separated floating electrodes with that of a sensor including one long floating electrode which has the same area as the total area of those seven separated floating electrodes. The results show that the sensor with a long floating electrode exhibited lower sensitivity than the sensor with seven separated floating electrodes even though both devices had the same surface area of floating electrodes. This result shows that the enhanced sensitivity of our sensors with larger number of floating electrodes resulted from the increased number of Schottky barriers formed at the floating electrodes rather than the increased surface area of the floating electrodes. Fig. 3d shows Langmuir isotherm curve fitting for the two kinds of DNA with different number of bases. In this experiment, the measurable concentration range of the sensor was much broader for 25-mer DNA than for 16-mer DNA. For 16-mer DNA, its hybridization process was detected from the 100 pM concentration. On the other hand, 25-mer target DNA can be detected from 100 fM concentration (Fig. S2 in Supporting Information). The equilibrium constants for those two DNAs were estimated as 5.2 × 107 M−1 (16mer DNA) and 1.8 × 108 M−1 (25-mer DNA), respectively. These values are similar to those reported previously [20–22]. This result clearly shows that the equilibrium constant (K) of DNA binding determines the sensitivity of the sensor device as predicted earlier. 4. Conclusions We successfully developed floating electrode-based sensors for the DNA detection with controllable responses. Here, we showed that sensor responses could be enhanced simply by increasing the number of floating electrodes. The enhanced sensor response can be explained by the increased number of Schottky barriers formed at the floating electrodes. The experimental results were analyzed using the Langmuir isotherm theory, and we confirmed that the sensor response could be controlled by the modulation of two main parameters, the characteristic parameter of sensor device (k) and an equilibrium constant between probe DNA and target DNA (K). As expected from the theoretical model, the sensor responses could be controlled simply by varying the number of floating electrodes, related to the parameter k. We expect our work can provide an efficient strategy to enhance the sensitivity and to control the response of biosensors.
[2] N. Kumar, A. Dorfman, J.I. Hahm, Ultrasensitive DNA sequence detection using nanoscale ZnO sensor arrays, Nanotechnology 17 (2006) 2875–2881. [3] T.I. Abdullin, I.I. Nikitina, D.G. Ishmukhametava, G.K. Budnikov, O.A. Konovalova, M.Kh. Salakhov, Carbon nanotube-modified electrodes for electrochemical DNA-sensors, Journal of Analytical Chemistry 62 (2007) 599–603. [4] T.G. Drummond, M.G. Hill, J.K. Barton, Electrochemical DNA sensors, Nature Biotechnology 21 (2003) 1192–1199. [5] E.L. Gui, L.J. Li, K. Zhang, Y. Xu, X. Dong, X. Ho, P.S. Lee, J. Kasim, Z.X. Shen, J.A. Rogers, S.G. Mhaisalkar, DNA sensing by field-effect transistors based on networks of carbon nanotubes, Journal of the American Chemical Society 129 (2007) 14427–14432. [6] P. He, L. Dai, Aligned carbon nanotube-DNA electrochemical sensors, Chemical Communications 46 (2004) 348–349. [7] P.H. Yeh, Z. Li, Z.L. Wang, Schottky-gated probe-free ZnO nanowire biosensor, Advanced Materials 21 (2009) 4975–4978. [8] H.R. Byon, H.C. Choi, Network single-walled carbon nanotube-field effect transistors (SWNT-FETs) with increased Schottky contact area for highly sensitive biosensor applications, Journal of the American Chemical Society 128 (2006) 2188–2189. [9] P.K. Basu, N. Saha, S.K. Jana, H. Saha, A. Lloyd Spetz, S. Basu, Schottky junction methane sensors using electrochemically grown nanocrystalline-nonporous ZnO thin films, Journal of Sensors (2009) 1–9, Article ID 790476. [10] E.D. Minot, A.M. Janssens, I. Heller, H.A. Heering, C. Dekker, S.G. Lemay, Carbon nanotube biosensors: the critical role of the reference electrode, Applied Physics Letters 91 (2007) 093507. [11] S. Heinze, J. Tersoff, R. Martel, V. Derycke, J. Appenzeller, Ph. Avouris, Carbon nanotubes as Schottky barrier transistors, Physical Review Letters 89 (2002) 106801–106805. [12] I. Heller, A.M. Janssens, J. Mannik, E.D. Minot, S.G. Lemay, C. Dekker, Identifying the mechanism of biosensing with carbon nanotube transistors, Nano Letters 8 (2008) 591–595. [13] R.J. Chen, H.C. Choi, S. Bangsaruntip, E. Yenilmez, X. Tang, Q. Wang, Y.-L. Chang, H. Dai, An investigation of the mechanisms of electronic sensing of protein adsorption on carbon nanotube devices, Journal of the American Chemical Society 126 (2004) 1563–1568. [14] X. Tang, S. Bansaruntip, N. Nakayama, E. Yenilmez, Y.I. Chang, Q. Wang, Carbon nanotube DNA sensor and sensing mechanism, Nano Letters 6 (2006) 1632–1636. [15] E.S. Snow, J.P. Novak, M.D. Lay, F.K. Perkins, 1/f noise in single-walled carbon nanotube devices, Applied Physics Letters 85 (2004) 4172–4174. [16] F.N. Hooge, T.G.M. Kleinpenning, L.K.J. Vandamme, Experimental studies on 1/f noise, Reports on Progress in Physics 44 (1981) 479–532. [17] P.G. Collins, M.S. Fuhrer, A. Zettl, 1/f noise in carbon nanotubes, Applied Physics Letters 76 (2000) 894–896. [18] J. Kang, J. Lee, T.H. Kim, J. Park, M.J. Sung, S. Hong, Large-scale assembly of carbon nanotube-based flexible circuits for DNA sensors, Nanotechnology 19 (2008) 135305. [19] T.H. Kim, J. Lee, S. Hong, Highly sensitive environmental nanosensors based on anomalous response of carbon nanotube conductance to mercury ions, Journal of Physical Chemistry C 113 (2009) 19393–19396. [20] Y.L. Bunimovich, Y.S. Shin, W.S. Yeo, M. Amori, G. Kwong, J.R. Heath, Quantitative real-time measurements of DNA hybridization with alkylated nonoxidized silicon nanowires in electrolytes solution, Journal of the American Chemical Society 128 (2006) 16323–16331. [21] F. Fixe, H.M. Branz, N. Louro, V. Chu, D.M.F. Prazeres, J.P. Conde, Electric field assisted immobilization and hybridization of DNA oligomers on thin-film microchips, Nanotechnology 16 (2005) 2061–2071. [22] Z. Zhang, W. Knoll, R. Forch, Amino-functionalized plasma polymer films for DNA immobilization and hybridization, Surface and Coatings Technology 200 (2005) 993–995.
Biographies Acknowledgments This research was supported by the Conversing Research Center Program (No. 2011K000683). S.H. acknowledges the support from the National Research Foundation of Korea (NRF) grant (No. 20110000390) and the Samsung electronics. Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.snb.2012.04.063. References [1] M. Hazani, R. Naaman, F. Hennrich, M.M. Kappes, Confocal fluorescence imaging of DNA-functionalized carbon nanotubes, Nano Letters 3 (2003) 153–155.
Byeongju Kim is a senior engineer at Samsung Electronics Company in Korea. He received PhD degree in physics at Seoul National University, Korea, in 2011. His research interests include nano-materials and biocompatible nanostructures. Joohyung Lee is a senior engineer at Samsung Electronics Company in Korea. He received PhD degree in physics at Seoul National University, Korea, in 2012. His research interests include nano-materials and biosensors. Seon Namgung received BS and PhD degrees in physics at Seoul National University, Korea, in 2011. His current interests include the biological applications of nanomaterials and biosensors. Moon-Sook Lee is a research staff in the Bio Research Cencer at Samsung Advanced Institute of Technology in Korea. She received BS degree at Korea University in Korea. Then, she got PhD degree in Physical Chemistry at New York University, USA, in 1999. Her research interests include the convergence of nano-device technology and molecular, cellular biology for biomedical applications. Jeongsu Kim received BS degree in physics at Chonnam National University, and now is a MS candidate in physics at Seoul National University in Korea. His current interests include nano-materials and the fabrication of nano devices.
B. Kim et al. / Sensors and Actuators B 169 (2012) 182–187 Jae Yeol Park is a professor in the Department of Automotive Engineering at Doowon Technical University College in Korea. He received BS, MS degree and PhD degree in material engineering at Seoul National University, Korea, in 2003. His current interests include nano-material device design and structural analysis.
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Seunghun Hong is a professor in the Department of Physics and Astronomy and the Department of Biophysics and Chemical Biology at Seoul National University in Korea. He received BS and MS degree at Seoul National University, in Korea. Then, he got PhD degree at Purdue University, USA, in 1998. His research interests include nanoscale hybrid systems comprised of solid state devices and organic materials.
Contents lists available at SciVerse ScienceDirect
Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb
DNA sensors based on CNT-FET with floating electrodes Byeongju Kim a , Joohyung Lee a , Seon Namgung a , Jeongsu Kim a , Jae Yeol Park b , Moon-Sook Lee c , Seunghun Hong a,d,∗ a
Department of Physics and Astronomy, Seoul National University, Seoul 151-747, Republic of Korea Department of Automotive Engineering, Doowon Technical University College, Anseong 456-718, Republic of Korea Bio Center, Samsung Advanced Institute of Technology, Yong-In 446-712, Republic of Korea d Department of Biophysics and Chemical Biology (WCU Program), Seoul National University, Seoul 151-747, Republic of Korea b c
a r t i c l e
i n f o
Article history: Received 24 December 2011 Received in revised form 14 April 2012 Accepted 22 April 2012 Available online 28 April 2012 Keywords: Carbon nanotube Sensor Floating electrode Schottky barrier DNA
a b s t r a c t We report a successful development of a floating electrode-based DNA sensor with controllable responses. Here, metallic floating electrodes were fabricated to form Schottky barriers between carbon nanotubes and the floating electrodes. We showed that the sensor response could be enhanced by increasing the number of floating electrodes. We also analyzed the response of the sensors based on the Langmuir isotherm theory. © 2012 Elsevier B.V. All rights reserved.
1. Introduction The electrical detection method of DNA can have several advantages over bulky fluorescence-based DNA detection method, such as label-free detection, high sensitivity, and easy integration with other devices for practical applications [1–6]. Recently, diverse DNA sensors based on nanomaterials have been developed. Among them, carbon nanotubes (CNT) network-based DNA sensors have attracted much attention due to its high sensitivity. In the CNT network-based DNA sensors, the modulation of Schottky barriers by DNA hybridization has been known to play a crucial role in detecting target DNA. When the target DNA was bound to probe DNA, the height of Schottky barriers at the contact regions between CNTs and metal electrodes were varied. Since the height of Schottky barriers controls the electric currents of the devices, the binding of target DNA could be detected by monitoring the change of the electric currents in the devices. There has been much effort to enhance the sensitivity of the Schottky barrier-based sensors by forming a nonsymmetrical Schottky contact [7], by increasing Schottky contact area using a shadow mask [8], and by varying the contact metal
∗ Corresponding author at: Department of Physics and Astronomy, Seoul National University, Seoul 151-747, Republic of Korea. Tel.: +82 2 880 1343; fax: +82 2 884 3002. E-mail address: [email protected] (S. Hong). 0925-4005/$ – see front matter © 2012 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.snb.2012.04.063
[9]. However, in the previous methods, it was difficult to predict or to control the sensor responses for practical applications. Herein, we developed a floating electrode-based sensor for the detection of DNA molecules with controllable responses. We showed that sensor responses could be enhanced simply by increasing the number of floating electrodes. The experimental results were explained by the modulation of Schottky-barriers formed at the floating electrodes. We also demonstrated that experimental results fitted well to the Langmuir isotherm theory. Our approach could be a simple but efficient way to improve the sensitivity of CNT network-based sensors and should provide an important insight regarding Schottky barrier-based sensors in general.
2. Experimental 2.1. Materials Purified single-walled carbon nanotubes (swCNTs) were purchased from Carbon Nanotechnologies Inc. (USA) and used as received. The swCNTs have a diameter of 0.7–2 nm and a length of 2–3 m. Mercaptohexanol (MCH), octadecyltrichlorosilane (OTS) and other chemical reagents were purchased from Sigma–Aldrich (USA) and used as received. DNAs used in these experiments were purchased from GenoTech Company (Korea). Their base sequences are listed below:
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16 mer probe DNA: Thiol-C6 -5 -GCC ACA AAC ACC ACA A-3 16 mer target DNA: 5 -TTG TGG TGT TTG TGG C-3 16 mer noncomplementary DNA: 5 -CAT GGT TGA TCC GTT C-3 25 mer probe DNA: Thiol-C6 -5 -AGA CCT CCA GTC TCC ATG TTA CGT C-3 25 mer target DNA: 5 -GAC GTA ACA TGG AGA CTG GAG GTC T-3 25 mer noncomplementary DNA: 5 -ACG CTG AGT ACG GGT GCA AGA GTC A-3 2.2. Fabrication of CNT-field effect transistor (CNT-FET) devices To create CNT channel patterns, we first patterned AZ 5214 photoresist (PR) on a silicon oxide wafer via standard photolithography. Then, the wafer was placed in an OTS solution (1:500, v/v in hexane) for 5 min. Subsequently, the PR patterns were removed using acetone. Then, the patterned silicon oxide wafer was placed in a CNT solution (concentration 0.1 mg/ml) for 30 s, which was prepared by dispersing CNTs in 1,2-dichlorobenzene with ultrasonication for 1 h. This step allowed CNTs to be adsorbed selectively onto bare SiO2 regions in the wafer, while the methyl-terminated OTS selfassembled monolayer (SAM) blocked non-specific adsorption of the CNTs. The wafer was rinsed thoroughly with 1,2-dichlorobenzene and dried with nitrogen gas. After then, metal electrodes (30 nm Au on 10 nm Ti) were formed via standard photolithography, thermal evaporation, and lift-off process. Finally, we covered the source and drain electrodes with a photoresist layer to eliminate leakage currents from the electrodes in a buffer solution. 2.3. Functionalization of electrodes with probe DNA First, the CNT-FETs were placed in a probe DNA solution (10 M in phosphate buffered saline (PBS, pH 7.4, Fisher Scientific)) for 10 h to functionalize the Au electrodes with the probe DNA. The devices were rinsed with PBS and dried with a N2 stream. The devices were placed in MCH solution (10 mM in deionized water) overnight and dried with a N2 stream. 2.4. Electrical detection of target DNA using CNT-FETs In DNA detection experiments, the source–drain currents of the floating electrode-based CNT-FET devices were measured by a semiconductor characterization system (Keithley, 4200, USA). The system was connected to a probe station that makes electrical contacts to the source and drain electrodes of the CNT-FET devices. During the measurement process, the source–drain (SD) bias (Vsd ) of 100 mV was applied. A PBS buffer solution was dropped onto the channel regions of the devices. Then, an Ag/AgCl electrode, used as a gate electrode, was inserted into the solution, and a gate voltage (Vlg ) of 0 V was applied. The Ag/AgCl electrode was used to minimize the side effects such as electrochemical reactions at the reference electrode–solution interface [10]. After the base level of the signal was stabilized, the liquid gate characteristics of the device were measured. For liquid gating measurements before DNA hybridization, the gate voltage was swept from −0.5 to +0.5 V with a Vsd of 100 mV. Then, for the detection of DNA hybridization, target DNA solutions with various DNA concentrations, from 100 fM to 10 M, were introduced into the buffer solution. 3. Results and discussion 3.1. Sensing mechanism Fig. 1a illustrates the structure of our floating electrode-based DNA sensor (upside) and its corresponding energy-band diagram (downside). Here, probe DNA was adsorbed only onto the surface of the floating electrodes. Source and drain electrodes were covered
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by photoresist to block leakage currents. The energy-band diagram shows the formation of Schottky barriers at the contact regions between CNTs and metal electrodes. These Schottky barriers determine the conductance of the device [11,12]. Fig. 1b shows the schematic diagram of the target DNA sensing experiment. When a target DNA solution was injected, the target DNA molecules were hybridized with their complementary probe DNA molecules on the floating electrodes (upside). During the hybridization process, the work function of the floating electrodes was decreased, and the Schottky barrier height was increased (downside). Since this increased Schottky barrier was acting as a potential barrier for hole-carriers, the currents in the channel region were decreased [13,14]. Note that the electrical conductance of device was determined solely by the floating electrodes because source and drain electrodes were passivated with PR. Therefore, the capability of the devices to detect target DNA could be controlled simply by varying the number and the shape of the floating electrodes. 3.2. Device layout and electrical characterization of CNT-FETs Fig. 2a shows the channel region of our floating electrode-based CNT-FET devices. The scanning electron microscopy (SEM) image (left) clearly shows that each floating electrode has dimensions of 12 m width and 2 m length. Here, the CNT network channel has dimensions of 2 m width and 40 m length. Note that the size and the number of floating electrodes can be readily controlled by simple photolithography processes. The right image of Fig. 2a is an atomic force microscope (AFM) image showing the CNT channel region between floating electrodes. Note that the CNTs were assembled only on a bare SiO2 region. The electrical properties of the floating electrode-based CNTFETs were investigated. Fig. 2b shows the typical liquid gating effect of our floating electrode-based CNT-FETs. The liquid gating effect was measured by applying a liquid gate bias (Vlg ) to the PBS buffer solution on a floating electrode-based CNT-FET. Here, an Ag/AgCl electrode was used as a liquid gate electrode, and a gate bias was swept from −0.5 V to 0.5 V with a source–drain bias of 0.1 V. The graph shows that the device also had p-type characteristics in the PBS buffer solution. Since the electric currents were drastically changed in the small range of voltage sweep, this device could be utilized as a sensitive sensor. We also analyzed the noise characteristics of our devices. The inset of Fig. 2c shows a typical noise spectrum of our devices with various numbers of floating electrodes at low frequency ranges. The graph shows a typical 1/f noise behavior, and the noise spectrum can be fitted by SI ∼ 1/fˇ with ˇ values of 1. The graph also shows that the noise amplitude was increased as the number of floating electrodes was increased. We confirmed that the noise amplitude A is linearly proportional to the device resistance R following the relationship of A ≈ 10−11 × R, which is a typical behavior of CNT networks (Fig. 2c) [15–17]. This result indicates that the noise of the floating electrode-based devices was dominated by CNT-networks in the channel region. 3.3. Detection of target DNA using floating electrode-based CNT-FETs Fig. 3a shows the response of our DNA sensors to the injection of non-complementary and complementary DNAs. When non-complementary DNA was injected, there were no significant changes in electric currents. However, when complementary DNA was injected, the significant decrease of electric currents was observed. The reduction of the electric currents can be explained by the hybridization of complementary target DNA molecules with the probe DNA molecules fixed on floating electrodes. The hybridization resulted in the increase of the height of Schottky barrier, which consequently reduced the currents of the devices (Fig. 1) [12,18].
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Fig. 1. Sensing mechanism of floating electrode-based CNT-FET sensors. (a) (Upside) Schematic diagram of a fabricated device before target DNA binding. Probe DNA was adhered onto a floating electrode surface, and source and drain electrodes were covered with photoresist to block leakage currents. (Downside) Energy-band diagram of the floating electrode-based CNT-FET device. Schottky barriers were formed at contact regions between CNTs and metal electrodes. Here, hole-carriers passed through the Schottky barriers formed at the floating electrode. (b) (Upside) Schematic diagram of target DNA sensing. The target DNA was hybridized with probe DNA on the surface of the floating electrode. (Downside) Energy-band diagram of the floating electrode-based CNT-FET device after target DNA binding. When the target DNA was hybridized with probe DNA, the work function of the floating electrode was decreased, which increased the height of Schottky barriers. Due to the increased height of Schottky barriers, the currents of the device, dominated by hole-carriers, were decreased.
Fig. 2. Characterization of floating electrode-based CNT-FETs. (a) SEM (left) and AFM (right) image of a floating electrode-based CNT-FET which has seven floating electrodes. Individual floating electrode had dimensions of 12 m width and 2 m length, while a CNT network channel had dimensions of 2 m width and 40 m length. (b) Liquid gating effect of a floating electrode-based CNT-FET. Here, an Ag/AgCl electrode was used as a liquid gate, and the gate bias was swept from −0.5 V to 0.5 V with a source–drain bias of 0.1 V. (c) Graph showing the noise amplitude (A) versus the channel resistance (R) relation for the floating electrode-based CNT-FETs with various numbers of floating electrodes. A fitting line represents A ≈ 10−11 × R. The insert shows the noise power spectrum (SI ) of the floating electrode-based CNT-FETs with various numbers of floating electrodes at low frequency ranges.
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Fig. 3. Response of a floating electrode-based CNT-FET sensor to the binding of target DNA to probe DNA. (a) Real-time response of a floating electrode-based DNA sensor to the injection of DNA. A drastic change of source–drain currents was observed only when complementary target DNA was introduced. Arrows indicate the points of DNA injections. (b) Real-time response of a floating electrode-based DNA sensor to the injection of target DNA with various concentrations. The significant reduction of currents was observed on the addition of target DNA from 1 nM to 10 M concentration. Arrows indicate the points of target DNA injections. (c) Response curves of floating electrodebased DNA sensors with various numbers of floating electrodes. Here, the sensitivity of floating electrode-based DNA sensors was defined as G/G0 . Note that the sensitivity of the sensors at certain concentration was increased as the number of floating electrodes was increased. (d) Response curves of floating electrode-based DNA sensors to 16-mer and 25-mer target DNAs. Note that the detection limit for 25-mer DNA was 100 fM, while that for 16-mer DNA was 100 pM.
This result clearly shows that our floating electrode-based CNTFETs could selectively detect the complementary target DNA in real time. When we functionalize floating electrodes by probe DNA molecules, the probe DNA molecules could wrap CNTs. These DNA molecules are known to play little role in hybridization with complementary DNA, while they block the non-specific adsorption of analytes [14]. Fig. 3b shows the real time measurement of the source–drain currents in our sensor device during the introduction of target DNA molecules with various concentrations ranging from 1 nM to 10 M. The addition of target DNA molecules caused a sharp decrease in the source–drain currents, and then the source–drain currents were gradually saturated. The change of currents (G/G0 ) increased as the concentration of target DNA solutions was elevated. The sensor response could be analyzed by a model based on the Langmuir isotherm theory [19]. If we assume that the maximum surface density of target DNA on the floating electrodes was Cs max and that target DNA molecules were hybridized with probe DNA molecules following the Langmuir isotherm, the surface density Cs of target DNA molecules bound to probe DNA molecules can be written as Cs =
Cs max · C 1/K + C
where C and K are the concentration of target DNA molecules in the solution and the equilibrium constant between probe DNA and target DNA, respectively. Furthermore, if we assume that the sensor signal (G/G0 ) can be approximated as |G/G0 | ∼ kCs where k is
a constant representing the characteristics of sensor device, the sensitivity of the sensor can be written as
Cs max · C G G = k 1/K + C 0
where G0 and G represent the original conductance of device and the conductance change of the sensor, respectively. From the above equation, we could deduce that the both parameters of k and K have great importance related to the sensitivity of the sensor device where k and K are related to the number of floating electrodes and the binding between target DNA and probe DNA, respectively. Fig. 3c shows the response of the sensor devices with various numbers of floating electrodes. Here, the sensitivity of floating electrode-based DNA sensors was defined as G/G0 . Using the above theoretical model, we performed curve fitting to experimental data sets with a large concentration range. The graph clearly shows that the experimental results fit well with the theoretical model. In previous works, we also applied the similar equation for the estimation of the equilibrium constant for the binding of mercury ions onto a CNT surface [19]. Here, all the equilibrium constants of probe DNA to target DNA with various numbers of floating electrodes had values in the same order (∼107 M−1 ). This result indicates that the equilibrium constant for given pair of DNA molecules did not change even though the structure of the sensors was varied. Note that the sensitivity of the sensors at a certain concentration was increased as the number of floating electrodes was increased. For example, the sensitivity of the sensor at 1 nM concentration with one floating electrode was 8%, while that with seven floating electrodes was 21%. As mentioned earlier, the
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sensor response was controllable by varying the number of floating electrodes. We performed control experiments to check if the enhanced sensitivity of our sensors with larger number of floating electrodes came simply from the increased surface area of the floating electrodes rather than the increased number of the floating electrodes (Fig. S1 in Supporting Information). Here, we compared the sensitivity of a sensor including seven separated floating electrodes with that of a sensor including one long floating electrode which has the same area as the total area of those seven separated floating electrodes. The results show that the sensor with a long floating electrode exhibited lower sensitivity than the sensor with seven separated floating electrodes even though both devices had the same surface area of floating electrodes. This result shows that the enhanced sensitivity of our sensors with larger number of floating electrodes resulted from the increased number of Schottky barriers formed at the floating electrodes rather than the increased surface area of the floating electrodes. Fig. 3d shows Langmuir isotherm curve fitting for the two kinds of DNA with different number of bases. In this experiment, the measurable concentration range of the sensor was much broader for 25-mer DNA than for 16-mer DNA. For 16-mer DNA, its hybridization process was detected from the 100 pM concentration. On the other hand, 25-mer target DNA can be detected from 100 fM concentration (Fig. S2 in Supporting Information). The equilibrium constants for those two DNAs were estimated as 5.2 × 107 M−1 (16mer DNA) and 1.8 × 108 M−1 (25-mer DNA), respectively. These values are similar to those reported previously [20–22]. This result clearly shows that the equilibrium constant (K) of DNA binding determines the sensitivity of the sensor device as predicted earlier. 4. Conclusions We successfully developed floating electrode-based sensors for the DNA detection with controllable responses. Here, we showed that sensor responses could be enhanced simply by increasing the number of floating electrodes. The enhanced sensor response can be explained by the increased number of Schottky barriers formed at the floating electrodes. The experimental results were analyzed using the Langmuir isotherm theory, and we confirmed that the sensor response could be controlled by the modulation of two main parameters, the characteristic parameter of sensor device (k) and an equilibrium constant between probe DNA and target DNA (K). As expected from the theoretical model, the sensor responses could be controlled simply by varying the number of floating electrodes, related to the parameter k. We expect our work can provide an efficient strategy to enhance the sensitivity and to control the response of biosensors.
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Biographies Acknowledgments This research was supported by the Conversing Research Center Program (No. 2011K000683). S.H. acknowledges the support from the National Research Foundation of Korea (NRF) grant (No. 20110000390) and the Samsung electronics. Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.snb.2012.04.063. References [1] M. Hazani, R. Naaman, F. Hennrich, M.M. Kappes, Confocal fluorescence imaging of DNA-functionalized carbon nanotubes, Nano Letters 3 (2003) 153–155.
Byeongju Kim is a senior engineer at Samsung Electronics Company in Korea. He received PhD degree in physics at Seoul National University, Korea, in 2011. His research interests include nano-materials and biocompatible nanostructures. Joohyung Lee is a senior engineer at Samsung Electronics Company in Korea. He received PhD degree in physics at Seoul National University, Korea, in 2012. His research interests include nano-materials and biosensors. Seon Namgung received BS and PhD degrees in physics at Seoul National University, Korea, in 2011. His current interests include the biological applications of nanomaterials and biosensors. Moon-Sook Lee is a research staff in the Bio Research Cencer at Samsung Advanced Institute of Technology in Korea. She received BS degree at Korea University in Korea. Then, she got PhD degree in Physical Chemistry at New York University, USA, in 1999. Her research interests include the convergence of nano-device technology and molecular, cellular biology for biomedical applications. Jeongsu Kim received BS degree in physics at Chonnam National University, and now is a MS candidate in physics at Seoul National University in Korea. His current interests include nano-materials and the fabrication of nano devices.
B. Kim et al. / Sensors and Actuators B 169 (2012) 182–187 Jae Yeol Park is a professor in the Department of Automotive Engineering at Doowon Technical University College in Korea. He received BS, MS degree and PhD degree in material engineering at Seoul National University, Korea, in 2003. His current interests include nano-material device design and structural analysis.
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Seunghun Hong is a professor in the Department of Physics and Astronomy and the Department of Biophysics and Chemical Biology at Seoul National University in Korea. He received BS and MS degree at Seoul National University, in Korea. Then, he got PhD degree at Purdue University, USA, in 1998. His research interests include nanoscale hybrid systems comprised of solid state devices and organic materials.