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A novel real-time immunoassay utilizing an electro-immunosensing microchip and gold nanoparticles for signal enhancement
Sensors and Actuators B 117 (2006) 451–456
A novel real-time immunoassay utilizing an electro-immunosensing microchip and gold nanoparticles for signal enhancement Min Li a , Yu-Cheng Lin a,∗ , Kai-Chun Su a , Yu-Tsung Wang a , Tsung Chain Chang b , Hong-Ping Lin c a
Department of Engineering Science, National Cheng Kung University, 1 University Road, Tainan 701, Taiwan, ROC b Department of Medical Laboratory Science and Biotechnology, 1 University Road, Tainan, Taiwan, ROC c Department of Chemistry, National Cheng Kung University, 1 University Road, Tainan, Taiwan, ROC Accepted 15 December 2005 Available online 17 February 2006
Abstract This study focused on the development of an electro-immunosensing (EIS) microchip for real-time measurement of antibody–antigen recognition in immunoassay. Instead of the enzyme being conjugated with a secondary antibody for detection, the process of electro-immunoassay was found to be relatively simpler compared to the conventional enzyme-linked immunosorbent assay (ELISA). This novel testing method was designed using the wave impedance theory to measure the antibody–antigen recognition. Based on the phase variety in the frequency domain, the detection limit of protein A was 1 ng/mL. It was found that the antibody–antigen recognition has an obvious peak of the phase angle near 1.61 GHz. The EIS chip had higher sensitivity and a shorter assay time than the ELISA. The sensitivity of the immunoassay on the EIS chip was 100-fold higher than that of conventional enzyme-linked immunosorbent assay (ELISA). Using antibodies labelled with 13 nm gold nanoparticles (ANPs), the detection sensitivity of protein A was increased to 0.1 ng/mL. In addition, the EIS chip was capable of detecting the kinematics of antibody–antigen binding. © 2006 Elsevier B.V. All rights reserved. Keywords: Immunosensing; Impedance; Protein A; Gold nanoparticles
1. Introduction Immunoassays are commonly used in clinical laboratories for the detection of a variety of antigens and antibodies. Basically, immunoassays are sensitive (1–10 ng/mL) and specific. The most commonly used format of immunoassays is ELISA, which is a very useful and powerful method for estimating antibody–antigen reaction [1–3]. The signal of the conventional enzyme-linked immunosorbent assay is produced by the enzyme conjugated to the secondary antibody. This way, the specific antigen will be detected by the special emission spectrum. However, ELISA has the disadvantages of using multiple incubations and washing steps. Researchers have tried to increase ELISA’s sensitivity and shorten the reaction time. Electric, optic, and magnetic technologies were all successively introduced to immunosensing [4–6]. The micro-electrode, used in electro-chemical detection,
can amplify a weak electric signal, thereby making it easy to increase the sensitivity of antibody–antigen recognition [7–9]. There are some studies that used the theory of complementary metal semiconductor (CMOS) to design a portable immunosensing unit [10–13]. The nanoparticles have many excellent physical properties, such as high surface/volume ratio, unique absorption spectrum and electric conductivity, which are beneficial and suitable for immunoassay material and methods. Gold colloid was used in the electrochemical method to increase the immobilization amount of antigen [14–17]. In this study, we used the MEMS technology and the electromagnetic wave reflection model to develop an EIS chip for immunoassay. In addition, the kinematics of the antigen– antibody reaction could be detected in real time as well. 2. Theory
∗
Corresponding author. Tel.: +886 6 276 2395; fax: +886 6 276 2329. E-mail address: [email protected] (Y.-C. Lin).
0925-4005/$ – see front matter © 2006 Elsevier B.V. All rights reserved. doi:10.1016/j.snb.2005.12.055
The antibody–antigen reaction measurement is similar to the thin film ones, and the difference in a variety of dielectric
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M. Li et al. / Sensors and Actuators B 117 (2006) 451–456
Fig. 1. Schematic drawing of the reflection coefficient measurement method.
properties (ρ) of the antibody–antigen layer will follow with or without antigen targeting. The volume of buffer is larger than that of the antibody–antigen layer. Based on the definition of bulk resistance (Eq. (1)), the impedance of the antibody–antigen layer (Ab/Ag layer) is higher than that of buffer and slide glass: R=
ρL wd
(1)
Fig. 2. Schematic drawing of an EIS chip: (A) the pattern of electrodes and (B) expanded view of the EIS chip.
Fig. 3. Illustrations of the assays for immunosensing: (A) EIS chip, (B) assay with using second IgG-Au nanoparticle complex in the sandwich format of the EIS chip, and (C) ELISA.
M. Li et al. / Sensors and Actuators B 117 (2006) 451–456
where the L and w are the space of both electrodes and the height of the layer (antibody layer dAb , antigen layer dAg , buffer db or glass dg ), respectively. Most of the current will go through the buffer and the slide glass (dg or db df ). As a result, the antibody–antigen reaction can hardly be measured by the LCR principle in low frequency. The impedance is measured in the high frequency domain, often by adopting the reflection coefficient measurement method. The reflection coefficient is the ratio of the reflected wave to the incident wave at the reflection point. The electrical mode for two planar microelectrodes is shown in Fig. 1. When a transverse electromagnetic wave (TEM waves) travels in the Ab/Ag layer, the energy of the wave is lost. The voltage reflection coefficient (ϕ) is correlated with the impedance, by Eq. (2) as follows: ϕ=
ZAg − ZAg−Ab ZAg + ZAg−Ab
453
photolithography and wet etching, were used to fabricate thinfilm micro-electrodes for the electrical probes. Gold and titanium thin films were thermally evaporated onto microscopic glass slides. Photolithography was used to transfer the electrode pattern to the thin films. The interdigitated electrodes were formed after chemical wet etching. The electrodes were made of 25 nm titanium and 500 nm gold, and their dimensions were 100 m
(2)
The antigen transmission coefficient and the antigen bodylabelling transmission coefficient are ZAg and ZAg−Ab respectively. When the non-specific antigen can label the IgG, the value of ZAg equalizes the value of ZAg−Ab and the ϕ equals to zero. The impedance of the antibody–antigen reaction is measured by the RF I–V method. The input channel signal represents the voltage across the Ab/Ag layer, and the output channel signal represents the current flowing through the Ab/Ag layer. The antibody and antigen form capacitive biolayers attached to the surface of the electrodes. An alternating input signal from the network analyzer with constant amplitude Vin at the electrode on the left side capacitor couples with the electrolyte solution, and continues to couple capacity with the electrode on the right side. Therefore, the output current Iout is strongly dependent on the total thickness of the biolayer and is given by Eq. (3): −1 Vin 2 2 2 Iout = ψ = Vin RB + + + ψ Ztotal iCAg iCAb iCB (3) where Ztotal is the total impedance of the system, is the angular frequency of the applied voltage, CAg , CAb , CB , RB are the capacitance of the antigen, the layer of immobilized antibodies, the buffer and the resistance of the buffer, respectively. As per the skin effect, the skin depth decreases as the frequency increases and the EIS chip becomes more sensitive. Based on the reflection coefficient theory and the skin effect, the impedance variety of the Ab/Ag layer can be distinguished by its own characteristic frequency. 3. Experimental The microelectrodes were designed to form a thin-film structure, resulting in a strong point discharge and better sensitivity [5–7]. The MEMS technology was used to fabricate the Au/Ti thin film electrodes. A 3 mm thick poly-dimethyl siloxane (PDMS) with a 6 mm square opening was formed from a predefined mold, and was bound to the surface of the glass with patterned electrodes to form the antibody–antigen reaction area [11]. Microfabrication processes, such as thermal evaporation,
Fig. 4. Detection limit of antigen: (A) the phase of the EIS chip in the frequency spectra (300 kHz to 1.8 GHz), (B) the impedance of the EIS chip in the frequency spectra (300 kHz to 1.8 GHz), (C) conventional ELISA.
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wide with 200 m space between electrodes. The schematic drawings of an EIS are shown in Fig. 2. Antigen protein A and anti-protein A-IgG-HRP (IgG) were selected to test the feasibility and performance of the EIS chip. Non-specific antigen beta-glucuronidase (GUD) was used as the control. Fig. 3 shows the schematic assay format of an electro-immunosensing assay and ELISA. The Spectrum and Impedance Measurement Analyzer, 4396B (Agilent Technologies, USA) using the RF I–V technique provided high-frequency range performance, extending up to 1.8 GHz. This technique was suitable to a frequency of 100 MHz and higher, and the scanning rate was 1 Hz per dot. To proceed with the conventional ELISA procedures, the first step was to add an aliqot of 50 L IgG solution into the reaction wells and place it in an incubator at 37 ◦ C for 1 h. In the second step, the wells were washed three times and blocked with phosphate buffered saline + 0.05% Tween 20 (PBST), and put in an incubator at 37 ◦ C for 1 h. The third step was to add 50 L of serially 10-fold diluted protein A (diluted in PBS containing 2% BSA) to the wells, followed by incubation at 37 ◦ C for 1 h.
The second washing step was then repeated. The fourth step was to add 50 L IgG in the wells and place it in an incubator at 37 ◦ C for 30 min. Finally, 40 L of 1 M P3 HO4 was added to the wells for stopping the reaction. Compared with the conventional ELISA procedures, which requires at least five steps, the EIS chip only needs three steps. It is evident that the immunoassay with the EIS chip is more convenient than the conventional ELISA. ANPs were prepared in an aqueous solution following the chemical reduction method. In the synthesizing process, an oil bath was used to reach the boiling temperature in the presence of hydrogen tetrachloraurate (HAuCl4 , 5 mM) and trisodium citrate dihydrate (1 mM). The Au3+ was reduced to its atom form by a reductant. The gold nanoparticles were covered with citrate anions, generating strong ionic force to prevent aggregation in a solution and bearing a net negative charge in a pH-adjusted solution. IgG should present a net positive charge in a solution at pH 6.0 due to the value of pI close to 7.0. Therefore, when the IgG (pH 6.0) solution was periodic gently mixed with the ANPs solution, IgG should become attached onto the surfaces of ANPs
Fig. 5. Specificity of the antibody–antigen reaction; (A) phase change after protein A binding to anti-protein A-IgG-HRP (IgG) in the electro-immunosensing, (B) phase change when GUD was used as a negative control antigen in electro-immunosensing, and (C) colorimetric detection of protein A using ELISA.
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through the attracting force of ionic interactions. The mixture of IgG (30 g/mL, 0.4 mL) and ANP (0.4 mL) was incubated at room temperature for 30 min, during which the IgG antibodies adsorbed onto the ANPs through a combination of ionic and hydrophobic interactions. After incubation for 30 min, the clear to pink supernatant obtained after centrifugation at 17,390 × g for 10 min of the ANP–IgG incubation solution consisted of unbound IgGs and was discarded. Finally, the resuspend solution (20 mM TBS with 0.1% BSA) was added to the remaining solution (a dark red, loosely packed sediment of the IgG-labeled ANP) to result in stable ANP–IgG solution. Conjugates can be stored at 4 ◦ C for more than 1 month without loss of activity. 4. Result and discussion Different concentrations of protein A, 0.1–1000 ng/mL, were used to test the sensitivity of the EIS chip. Compared with the chip coated only with IgG, the phase had an apparent variation at 1.61 GHz. The detection limit was 1 ng/mL in the phase chart, as shown in Fig. 4(A), similar to the impedance curve where antigen protein A at a concentration of 1 ng/mL can still be detected (Fig. 4(B)). However, the sensitivity of a conventional ELISA assay was around 10 ng/mL (Fig. 4(C)). When the IgG concentration was higher than 1000 ng/mL, the characteristic frequency peak, at 1.61 GHz, would be disappeared. The antigen protein A and its antibody IgG were immobilized on the chip when the EIS chip was used for testing. Beta-glucuronidase was used for specificity test. The impedance, i.e. phase, resistance, etc., was measured in a frequency range from 200 kHz to 1.8 GHz. The surface of the chip with IgG coating has an obvious peak of the phase angle near 1.61 GHz. When protein A was bound to the IgG, the phase peak decreased, as shown in Fig. 5(A). However, when GUD was used as the test antigen, the phase peak was still found at 1.61 GHz, which was the same peak frequency of the IgG, as shown in Fig. 5(B). The results of the ELISA on a chip are shown in Fig. 5(C). As expected, the detection limit was 10 ng/mL of protein A, and no cross-reaction was found for GUD. This result proved that the ability of the EIS chip to detect an antigen–antibody reaction was similar to that of the conventional ELISA. When the ANPs-labeled secondary antibody was added to form the sandwich format (0.1 ng/mL protein A) on the EIS chip surface, the peak of the impedance signal was shifted from 1.6 GHz to 1.0 GHz and amplified apparently. Using the ANPs-labeled secondary antibody showed a better colormetric discrimination than that of protein A only, as shown in Fig. 6(A). In addition, the signal of ANPs showed better results (Fig. 6(B)). The phase peak was shifted to 1.0 GHz and successfully elevated by near 120◦ . For the conventional ELISA, the blocking protein (BSA) is often used to avoid non-specific antigen binding on the unoccupied solid phase, but it decreases the EIS chip sensitivity. As a result, the blocking protein was not used in the realtime electric-immunoassay experiment. The phase in real time detection (Fig. 7) was different from that in Fig. 5(A and B). Two phase peaks could be identified, a 0.8 GHz negative peak and a 1.17 GHz positive peak. Real-time measurement of the
Fig. 6. Signal enhancement using the ANPs in the electro-immunoassay: (A) 50 L of secondary antibody labeled with 13 nm ANPs added in the well, (B) the phase expression of the antibody–antigen recognition. The dotted square was the original detection peak (1.6 GHz) without using ANPs.
protein A–IgG interaction is shown in Fig. 7. The peaks of the phase gradually decreased to zero as per the various reaction times of protein A reacting with the IgG. The kinematics of the interaction of antibody–antigen were measured in real time. We demonstrated the ANPs that could increase the sensitivity of a EIS chip. Even though at the lowest concentration of protein A (0.1 ng/mL), the kinematics of the interaction of antibody–antigen could still be measured with ANPs. The distinguishable signals of impedance and phase were expressed between 0.82 GHz and 1.3 GHz during the binding of the antigen, protein A, to the antibody, IgG, as shown in Fig. 8.
Fig. 7. Real-time phase measurement of the antibody (IgG)–antigen (protein A) binding. The peaks of the phase at 0.8 GHz and 1.17 GHz were gradually approaching to 0◦ after the antigen was added, and the antigen bound to the antibody within 140 min.
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Taiwan, ROC for access to their equipment and for their technical support. Funding from the Ministry of Education and the National Science Council of Taiwan, ROC is gratefully acknowledged (NSC94-2323-B-006-005). References
Fig. 8. The enhancing effect of the ANP in the kinematics of the interaction of antibody–antigen (0.1 ng/mL), (A) the impedance expression of the antibody–antigen binding and (B) the phase expression of the antibody–antigen recognition.
5. Conclusions In this study, we successfully developed the EIS chip as a powerful sensor for immunoassay. We found that frequencies higher than 1 GHz could be detected according to the wave impedance theory and experimental results. In addition, the EIS chip was proved to have good specificity, and the assay sensitivity could reach 1 ng/mL protein A concentration. The detection sensitivity of the EIS chip could be amplified by 13 nm Au nanoparticles and reached 0.1 ng/mL. Compared to the conventional ELISA procedure, the reaction time with the EIS chip was shortened by one-third. Furthermore, the antigen–antibody binding process could be observed in real time. The experimental results proved that the electro-immunosensing microchip was able to simplify the procedure of immunoassay and increase its sensitivity, especially in real-time impedance measurement. Acknowledgements The authors would like to thank the Center for Micro/Nano Technology Research, National Cheng Kung University, Tainan,
[1] E. Olaussen, P. Rekvig, Screening tests for antinuclear antibodies (ANA): selective use of central nuclear antigens as a rational basis for screening by ELISA, J. Autoimmun. 13 (1999) 95–97. [2] C.J.M. Bartels, C. van Maanen, A.M. van der Meulen, T. Dijkstra, W. Wouda, Evaluation of three enzyme-linked immunosorbent assays for detection of antibodies to Neospora caninum in bulk milk, Vet. Parasitol. 131 (2005) 235–246. [3] G. Bruun, G. Halff, A. Nikaein, Comparison of panel reactive antibodies (PRA) by anti-human globulin (AHG) versus ELISA techniques, Hum. Immunol. 49 (1996) 107. [4] W. Laureyn, F. Frederix, P. Van Gerwen, G. Maes, Nanoscaled interdigitated gold electrodes for impedimetric immunosensing, in: Transducers’99, Sendai, Japan, 1999, pp. 884–886. [5] M.F. Kramer, D.V. Lim, A rapid and automated fiber optic-based biosensor assay for the detection of Salmonella in spent irrigation water used in the sprouting of sprout seeds, J. Food Prot. 67 (2004) 46–49. [6] S. Suzuki, T. Yammanashi, S.I. Tazawa, O. Kuroswa, M. Washizu, Quantitative analysis of DNA orientation in stationary ac electric field using fluorescence anisotropy, IEEE Solid-State Sens. Acturtors (1995) 1374–1382. [7] S.O. Kasap, Principles of Electronic Materials and Devices, second ed., McGraw-Hill, 2002, pp. 148–152. [8] K.R. Wehmeyer, H.B. Halsali, W.R. Heineman, Electrochemical investigation of hapten–antibody interactions by differential pulse polarography, Clin. Chem. 28 (1982) 1968–1972. [9] A.G. Gehring, S.I. Tu, Enzyme-linked immunomagnetic electrochemical detection of live Escherichia coli O157:H7 in apple juice, J. Food Prot. 68 (2005) 146–149. [10] P. Caillat, M. Belleville, C. Massit, Active CMOS biochip: an electroaddressed DNA probe, IEEE Int. Solid-State Circuit (1998) 272–273. [11] P.N. Gilles, D.J. Wu, C.B. Foster, P.J. Dillon, S.J. Chanock, Single nucleotide polymorphic discrimination by an electronic dot blot assay on semiconductor microchips, Nat. Biotechnol. 17 (1998) 365–370. [12] P. Swanson, R. Gelbart, S. Graham, L. Yang, W. Butler, D.E. Ackley, E. Sheldon, A fully multiplexed CMOS biochip for DNA analysis, in: Conference on Solid-State Sensors and Actuators, 1999, pp. 133–136. [13] G. Marrazza, I. Chianella, M. Mascini, Disposable DNA electrochemical sensor for hybridization detection, Biosens. Bioelectron. 14 (1999) 43–51. [14] M. Wang, L. Wang, G. Wang, X. Ji, Y. Bai, T. Li, S. Gong, J. Li, Application of impedance spectroscopy for monitoring colloid Au-enhanced antibody immobilization and antibody–antigen reactions, Biosens. Bioelectron. 19 (2004) 575–582. [15] F. Yingzi, Y. Ruo, T. Dianping, C. Yaqin, X. Lan, Study on the immobilization of anti-IgG on Au-colloid modified modified gold electrode via potentimetric immunosensor, cyclic voltammetry, and electrochemical impedance techniques, Colloid Surf. B Biointerfaces 40 (2005) 61–66. [16] S. Kubitschko, J. Spinke, T. Bruckner, S. Pohl, N. Oranth, Sensitivity enhancment of optical immunosensors with nanoparticles, Anal. Biochem. 253 (1997) 112–122. [17] C. Xia, F. Xin, C. Ke, G.L. Shen, R.Q. Yu, An electrochemical stripping metalloimmunoassay based on silver-enhanced gold nanoparticle label, Biosens. Bioelectron. 20 (2005) 1805–1812.
A novel real-time immunoassay utilizing an electro-immunosensing microchip and gold nanoparticles for signal enhancement Min Li a , Yu-Cheng Lin a,∗ , Kai-Chun Su a , Yu-Tsung Wang a , Tsung Chain Chang b , Hong-Ping Lin c a
Department of Engineering Science, National Cheng Kung University, 1 University Road, Tainan 701, Taiwan, ROC b Department of Medical Laboratory Science and Biotechnology, 1 University Road, Tainan, Taiwan, ROC c Department of Chemistry, National Cheng Kung University, 1 University Road, Tainan, Taiwan, ROC Accepted 15 December 2005 Available online 17 February 2006
Abstract This study focused on the development of an electro-immunosensing (EIS) microchip for real-time measurement of antibody–antigen recognition in immunoassay. Instead of the enzyme being conjugated with a secondary antibody for detection, the process of electro-immunoassay was found to be relatively simpler compared to the conventional enzyme-linked immunosorbent assay (ELISA). This novel testing method was designed using the wave impedance theory to measure the antibody–antigen recognition. Based on the phase variety in the frequency domain, the detection limit of protein A was 1 ng/mL. It was found that the antibody–antigen recognition has an obvious peak of the phase angle near 1.61 GHz. The EIS chip had higher sensitivity and a shorter assay time than the ELISA. The sensitivity of the immunoassay on the EIS chip was 100-fold higher than that of conventional enzyme-linked immunosorbent assay (ELISA). Using antibodies labelled with 13 nm gold nanoparticles (ANPs), the detection sensitivity of protein A was increased to 0.1 ng/mL. In addition, the EIS chip was capable of detecting the kinematics of antibody–antigen binding. © 2006 Elsevier B.V. All rights reserved. Keywords: Immunosensing; Impedance; Protein A; Gold nanoparticles
1. Introduction Immunoassays are commonly used in clinical laboratories for the detection of a variety of antigens and antibodies. Basically, immunoassays are sensitive (1–10 ng/mL) and specific. The most commonly used format of immunoassays is ELISA, which is a very useful and powerful method for estimating antibody–antigen reaction [1–3]. The signal of the conventional enzyme-linked immunosorbent assay is produced by the enzyme conjugated to the secondary antibody. This way, the specific antigen will be detected by the special emission spectrum. However, ELISA has the disadvantages of using multiple incubations and washing steps. Researchers have tried to increase ELISA’s sensitivity and shorten the reaction time. Electric, optic, and magnetic technologies were all successively introduced to immunosensing [4–6]. The micro-electrode, used in electro-chemical detection,
can amplify a weak electric signal, thereby making it easy to increase the sensitivity of antibody–antigen recognition [7–9]. There are some studies that used the theory of complementary metal semiconductor (CMOS) to design a portable immunosensing unit [10–13]. The nanoparticles have many excellent physical properties, such as high surface/volume ratio, unique absorption spectrum and electric conductivity, which are beneficial and suitable for immunoassay material and methods. Gold colloid was used in the electrochemical method to increase the immobilization amount of antigen [14–17]. In this study, we used the MEMS technology and the electromagnetic wave reflection model to develop an EIS chip for immunoassay. In addition, the kinematics of the antigen– antibody reaction could be detected in real time as well. 2. Theory
∗
Corresponding author. Tel.: +886 6 276 2395; fax: +886 6 276 2329. E-mail address: [email protected] (Y.-C. Lin).
0925-4005/$ – see front matter © 2006 Elsevier B.V. All rights reserved. doi:10.1016/j.snb.2005.12.055
The antibody–antigen reaction measurement is similar to the thin film ones, and the difference in a variety of dielectric
452
M. Li et al. / Sensors and Actuators B 117 (2006) 451–456
Fig. 1. Schematic drawing of the reflection coefficient measurement method.
properties (ρ) of the antibody–antigen layer will follow with or without antigen targeting. The volume of buffer is larger than that of the antibody–antigen layer. Based on the definition of bulk resistance (Eq. (1)), the impedance of the antibody–antigen layer (Ab/Ag layer) is higher than that of buffer and slide glass: R=
ρL wd
(1)
Fig. 2. Schematic drawing of an EIS chip: (A) the pattern of electrodes and (B) expanded view of the EIS chip.
Fig. 3. Illustrations of the assays for immunosensing: (A) EIS chip, (B) assay with using second IgG-Au nanoparticle complex in the sandwich format of the EIS chip, and (C) ELISA.
M. Li et al. / Sensors and Actuators B 117 (2006) 451–456
where the L and w are the space of both electrodes and the height of the layer (antibody layer dAb , antigen layer dAg , buffer db or glass dg ), respectively. Most of the current will go through the buffer and the slide glass (dg or db df ). As a result, the antibody–antigen reaction can hardly be measured by the LCR principle in low frequency. The impedance is measured in the high frequency domain, often by adopting the reflection coefficient measurement method. The reflection coefficient is the ratio of the reflected wave to the incident wave at the reflection point. The electrical mode for two planar microelectrodes is shown in Fig. 1. When a transverse electromagnetic wave (TEM waves) travels in the Ab/Ag layer, the energy of the wave is lost. The voltage reflection coefficient (ϕ) is correlated with the impedance, by Eq. (2) as follows: ϕ=
ZAg − ZAg−Ab ZAg + ZAg−Ab
453
photolithography and wet etching, were used to fabricate thinfilm micro-electrodes for the electrical probes. Gold and titanium thin films were thermally evaporated onto microscopic glass slides. Photolithography was used to transfer the electrode pattern to the thin films. The interdigitated electrodes were formed after chemical wet etching. The electrodes were made of 25 nm titanium and 500 nm gold, and their dimensions were 100 m
(2)
The antigen transmission coefficient and the antigen bodylabelling transmission coefficient are ZAg and ZAg−Ab respectively. When the non-specific antigen can label the IgG, the value of ZAg equalizes the value of ZAg−Ab and the ϕ equals to zero. The impedance of the antibody–antigen reaction is measured by the RF I–V method. The input channel signal represents the voltage across the Ab/Ag layer, and the output channel signal represents the current flowing through the Ab/Ag layer. The antibody and antigen form capacitive biolayers attached to the surface of the electrodes. An alternating input signal from the network analyzer with constant amplitude Vin at the electrode on the left side capacitor couples with the electrolyte solution, and continues to couple capacity with the electrode on the right side. Therefore, the output current Iout is strongly dependent on the total thickness of the biolayer and is given by Eq. (3): −1 Vin 2 2 2 Iout = ψ = Vin RB + + + ψ Ztotal iCAg iCAb iCB (3) where Ztotal is the total impedance of the system, is the angular frequency of the applied voltage, CAg , CAb , CB , RB are the capacitance of the antigen, the layer of immobilized antibodies, the buffer and the resistance of the buffer, respectively. As per the skin effect, the skin depth decreases as the frequency increases and the EIS chip becomes more sensitive. Based on the reflection coefficient theory and the skin effect, the impedance variety of the Ab/Ag layer can be distinguished by its own characteristic frequency. 3. Experimental The microelectrodes were designed to form a thin-film structure, resulting in a strong point discharge and better sensitivity [5–7]. The MEMS technology was used to fabricate the Au/Ti thin film electrodes. A 3 mm thick poly-dimethyl siloxane (PDMS) with a 6 mm square opening was formed from a predefined mold, and was bound to the surface of the glass with patterned electrodes to form the antibody–antigen reaction area [11]. Microfabrication processes, such as thermal evaporation,
Fig. 4. Detection limit of antigen: (A) the phase of the EIS chip in the frequency spectra (300 kHz to 1.8 GHz), (B) the impedance of the EIS chip in the frequency spectra (300 kHz to 1.8 GHz), (C) conventional ELISA.
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wide with 200 m space between electrodes. The schematic drawings of an EIS are shown in Fig. 2. Antigen protein A and anti-protein A-IgG-HRP (IgG) were selected to test the feasibility and performance of the EIS chip. Non-specific antigen beta-glucuronidase (GUD) was used as the control. Fig. 3 shows the schematic assay format of an electro-immunosensing assay and ELISA. The Spectrum and Impedance Measurement Analyzer, 4396B (Agilent Technologies, USA) using the RF I–V technique provided high-frequency range performance, extending up to 1.8 GHz. This technique was suitable to a frequency of 100 MHz and higher, and the scanning rate was 1 Hz per dot. To proceed with the conventional ELISA procedures, the first step was to add an aliqot of 50 L IgG solution into the reaction wells and place it in an incubator at 37 ◦ C for 1 h. In the second step, the wells were washed three times and blocked with phosphate buffered saline + 0.05% Tween 20 (PBST), and put in an incubator at 37 ◦ C for 1 h. The third step was to add 50 L of serially 10-fold diluted protein A (diluted in PBS containing 2% BSA) to the wells, followed by incubation at 37 ◦ C for 1 h.
The second washing step was then repeated. The fourth step was to add 50 L IgG in the wells and place it in an incubator at 37 ◦ C for 30 min. Finally, 40 L of 1 M P3 HO4 was added to the wells for stopping the reaction. Compared with the conventional ELISA procedures, which requires at least five steps, the EIS chip only needs three steps. It is evident that the immunoassay with the EIS chip is more convenient than the conventional ELISA. ANPs were prepared in an aqueous solution following the chemical reduction method. In the synthesizing process, an oil bath was used to reach the boiling temperature in the presence of hydrogen tetrachloraurate (HAuCl4 , 5 mM) and trisodium citrate dihydrate (1 mM). The Au3+ was reduced to its atom form by a reductant. The gold nanoparticles were covered with citrate anions, generating strong ionic force to prevent aggregation in a solution and bearing a net negative charge in a pH-adjusted solution. IgG should present a net positive charge in a solution at pH 6.0 due to the value of pI close to 7.0. Therefore, when the IgG (pH 6.0) solution was periodic gently mixed with the ANPs solution, IgG should become attached onto the surfaces of ANPs
Fig. 5. Specificity of the antibody–antigen reaction; (A) phase change after protein A binding to anti-protein A-IgG-HRP (IgG) in the electro-immunosensing, (B) phase change when GUD was used as a negative control antigen in electro-immunosensing, and (C) colorimetric detection of protein A using ELISA.
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through the attracting force of ionic interactions. The mixture of IgG (30 g/mL, 0.4 mL) and ANP (0.4 mL) was incubated at room temperature for 30 min, during which the IgG antibodies adsorbed onto the ANPs through a combination of ionic and hydrophobic interactions. After incubation for 30 min, the clear to pink supernatant obtained after centrifugation at 17,390 × g for 10 min of the ANP–IgG incubation solution consisted of unbound IgGs and was discarded. Finally, the resuspend solution (20 mM TBS with 0.1% BSA) was added to the remaining solution (a dark red, loosely packed sediment of the IgG-labeled ANP) to result in stable ANP–IgG solution. Conjugates can be stored at 4 ◦ C for more than 1 month without loss of activity. 4. Result and discussion Different concentrations of protein A, 0.1–1000 ng/mL, were used to test the sensitivity of the EIS chip. Compared with the chip coated only with IgG, the phase had an apparent variation at 1.61 GHz. The detection limit was 1 ng/mL in the phase chart, as shown in Fig. 4(A), similar to the impedance curve where antigen protein A at a concentration of 1 ng/mL can still be detected (Fig. 4(B)). However, the sensitivity of a conventional ELISA assay was around 10 ng/mL (Fig. 4(C)). When the IgG concentration was higher than 1000 ng/mL, the characteristic frequency peak, at 1.61 GHz, would be disappeared. The antigen protein A and its antibody IgG were immobilized on the chip when the EIS chip was used for testing. Beta-glucuronidase was used for specificity test. The impedance, i.e. phase, resistance, etc., was measured in a frequency range from 200 kHz to 1.8 GHz. The surface of the chip with IgG coating has an obvious peak of the phase angle near 1.61 GHz. When protein A was bound to the IgG, the phase peak decreased, as shown in Fig. 5(A). However, when GUD was used as the test antigen, the phase peak was still found at 1.61 GHz, which was the same peak frequency of the IgG, as shown in Fig. 5(B). The results of the ELISA on a chip are shown in Fig. 5(C). As expected, the detection limit was 10 ng/mL of protein A, and no cross-reaction was found for GUD. This result proved that the ability of the EIS chip to detect an antigen–antibody reaction was similar to that of the conventional ELISA. When the ANPs-labeled secondary antibody was added to form the sandwich format (0.1 ng/mL protein A) on the EIS chip surface, the peak of the impedance signal was shifted from 1.6 GHz to 1.0 GHz and amplified apparently. Using the ANPs-labeled secondary antibody showed a better colormetric discrimination than that of protein A only, as shown in Fig. 6(A). In addition, the signal of ANPs showed better results (Fig. 6(B)). The phase peak was shifted to 1.0 GHz and successfully elevated by near 120◦ . For the conventional ELISA, the blocking protein (BSA) is often used to avoid non-specific antigen binding on the unoccupied solid phase, but it decreases the EIS chip sensitivity. As a result, the blocking protein was not used in the realtime electric-immunoassay experiment. The phase in real time detection (Fig. 7) was different from that in Fig. 5(A and B). Two phase peaks could be identified, a 0.8 GHz negative peak and a 1.17 GHz positive peak. Real-time measurement of the
Fig. 6. Signal enhancement using the ANPs in the electro-immunoassay: (A) 50 L of secondary antibody labeled with 13 nm ANPs added in the well, (B) the phase expression of the antibody–antigen recognition. The dotted square was the original detection peak (1.6 GHz) without using ANPs.
protein A–IgG interaction is shown in Fig. 7. The peaks of the phase gradually decreased to zero as per the various reaction times of protein A reacting with the IgG. The kinematics of the interaction of antibody–antigen were measured in real time. We demonstrated the ANPs that could increase the sensitivity of a EIS chip. Even though at the lowest concentration of protein A (0.1 ng/mL), the kinematics of the interaction of antibody–antigen could still be measured with ANPs. The distinguishable signals of impedance and phase were expressed between 0.82 GHz and 1.3 GHz during the binding of the antigen, protein A, to the antibody, IgG, as shown in Fig. 8.
Fig. 7. Real-time phase measurement of the antibody (IgG)–antigen (protein A) binding. The peaks of the phase at 0.8 GHz and 1.17 GHz were gradually approaching to 0◦ after the antigen was added, and the antigen bound to the antibody within 140 min.
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Taiwan, ROC for access to their equipment and for their technical support. Funding from the Ministry of Education and the National Science Council of Taiwan, ROC is gratefully acknowledged (NSC94-2323-B-006-005). References
Fig. 8. The enhancing effect of the ANP in the kinematics of the interaction of antibody–antigen (0.1 ng/mL), (A) the impedance expression of the antibody–antigen binding and (B) the phase expression of the antibody–antigen recognition.
5. Conclusions In this study, we successfully developed the EIS chip as a powerful sensor for immunoassay. We found that frequencies higher than 1 GHz could be detected according to the wave impedance theory and experimental results. In addition, the EIS chip was proved to have good specificity, and the assay sensitivity could reach 1 ng/mL protein A concentration. The detection sensitivity of the EIS chip could be amplified by 13 nm Au nanoparticles and reached 0.1 ng/mL. Compared to the conventional ELISA procedure, the reaction time with the EIS chip was shortened by one-third. Furthermore, the antigen–antibody binding process could be observed in real time. The experimental results proved that the electro-immunosensing microchip was able to simplify the procedure of immunoassay and increase its sensitivity, especially in real-time impedance measurement. Acknowledgements The authors would like to thank the Center for Micro/Nano Technology Research, National Cheng Kung University, Tainan,
[1] E. Olaussen, P. Rekvig, Screening tests for antinuclear antibodies (ANA): selective use of central nuclear antigens as a rational basis for screening by ELISA, J. Autoimmun. 13 (1999) 95–97. [2] C.J.M. Bartels, C. van Maanen, A.M. van der Meulen, T. Dijkstra, W. Wouda, Evaluation of three enzyme-linked immunosorbent assays for detection of antibodies to Neospora caninum in bulk milk, Vet. Parasitol. 131 (2005) 235–246. [3] G. Bruun, G. Halff, A. Nikaein, Comparison of panel reactive antibodies (PRA) by anti-human globulin (AHG) versus ELISA techniques, Hum. Immunol. 49 (1996) 107. [4] W. Laureyn, F. Frederix, P. Van Gerwen, G. Maes, Nanoscaled interdigitated gold electrodes for impedimetric immunosensing, in: Transducers’99, Sendai, Japan, 1999, pp. 884–886. [5] M.F. Kramer, D.V. Lim, A rapid and automated fiber optic-based biosensor assay for the detection of Salmonella in spent irrigation water used in the sprouting of sprout seeds, J. Food Prot. 67 (2004) 46–49. [6] S. Suzuki, T. Yammanashi, S.I. Tazawa, O. Kuroswa, M. Washizu, Quantitative analysis of DNA orientation in stationary ac electric field using fluorescence anisotropy, IEEE Solid-State Sens. Acturtors (1995) 1374–1382. [7] S.O. Kasap, Principles of Electronic Materials and Devices, second ed., McGraw-Hill, 2002, pp. 148–152. [8] K.R. Wehmeyer, H.B. Halsali, W.R. Heineman, Electrochemical investigation of hapten–antibody interactions by differential pulse polarography, Clin. Chem. 28 (1982) 1968–1972. [9] A.G. Gehring, S.I. Tu, Enzyme-linked immunomagnetic electrochemical detection of live Escherichia coli O157:H7 in apple juice, J. Food Prot. 68 (2005) 146–149. [10] P. Caillat, M. Belleville, C. Massit, Active CMOS biochip: an electroaddressed DNA probe, IEEE Int. Solid-State Circuit (1998) 272–273. [11] P.N. Gilles, D.J. Wu, C.B. Foster, P.J. Dillon, S.J. Chanock, Single nucleotide polymorphic discrimination by an electronic dot blot assay on semiconductor microchips, Nat. Biotechnol. 17 (1998) 365–370. [12] P. Swanson, R. Gelbart, S. Graham, L. Yang, W. Butler, D.E. Ackley, E. Sheldon, A fully multiplexed CMOS biochip for DNA analysis, in: Conference on Solid-State Sensors and Actuators, 1999, pp. 133–136. [13] G. Marrazza, I. Chianella, M. Mascini, Disposable DNA electrochemical sensor for hybridization detection, Biosens. Bioelectron. 14 (1999) 43–51. [14] M. Wang, L. Wang, G. Wang, X. Ji, Y. Bai, T. Li, S. Gong, J. Li, Application of impedance spectroscopy for monitoring colloid Au-enhanced antibody immobilization and antibody–antigen reactions, Biosens. Bioelectron. 19 (2004) 575–582. [15] F. Yingzi, Y. Ruo, T. Dianping, C. Yaqin, X. Lan, Study on the immobilization of anti-IgG on Au-colloid modified modified gold electrode via potentimetric immunosensor, cyclic voltammetry, and electrochemical impedance techniques, Colloid Surf. B Biointerfaces 40 (2005) 61–66. [16] S. Kubitschko, J. Spinke, T. Bruckner, S. Pohl, N. Oranth, Sensitivity enhancment of optical immunosensors with nanoparticles, Anal. Biochem. 253 (1997) 112–122. [17] C. Xia, F. Xin, C. Ke, G.L. Shen, R.Q. Yu, An electrochemical stripping metalloimmunoassay based on silver-enhanced gold nanoparticle label, Biosens. Bioelectron. 20 (2005) 1805–1812.