Development of an AC electrokinetics-based immunoassay system for on-site serodiagnosis of infectious diseases

Sensors and Actuators A 171 (2011) 406–413

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Sensors and Actuators A: Physical journal homepage: www.elsevier.com/locate/sna

Development of an AC electrokinetics-based immunoassay system for on-site serodiagnosis of infectious diseases Xiaozhu Liu a,b,1 , Kai Yang a,1 , Ashutosh Wadhwa c , Shigetoshi Eda c , Shanshan Li a,d , Jie Wu a,∗ a

Department of Electrical Engineering and Computer Science, The University of Tennessee, Knoxville, TN 37996, USA Key Laboratory of Optoelectronic Technology and System of the Education Ministry of China, Chongqing University, Chongqing 400044, China Department of Forestry, Wildlife and Fisheries, The University of Tennessee, Knoxville, TN 37996, USA d School of Mechatronics Engineering, Harbin Institute of Technology, Harbin 150001, China b c

a r t i c l e

i n f o

Article history: Received 25 April 2011 Received in revised form 26 July 2011 Accepted 9 August 2011 Available online 22 August 2011 Keywords: AC electrokinetic Fluorescence detection Immunoassay Microfluidic chip Point-of-care

a b s t r a c t This paper presents a lab-chip immunoassay system that is based on AC electrothermal effect. It uses a poly (dimethylsiloxane) based microfluidic cartridge as the disposable immunoreactor, on which key processes such as reagent delivery, incubation and washing are performed using low voltage AC signals. A low cost, reconfigurable detection module is constructed based on light emitting diode induced fluorescence. Along with the employed AC electrothermal effect, the developed diagnostic system has demonstrated a much shorter incubation time than conventional pressure driven flow system. A tenfold acceleration in detection is achieved while safely differentiating between the positive and negative primary antibodies with large margin. Other merits of this immunoassay system include portability, rapid detection and low reagent consumption. © 2011 Elsevier B.V. All rights reserved.

1. Introduction Conventional immunoassay test is a multistage, labor-intensive process that usually consists of multiple reagents dispensing, incubations and several washing processes [1]. It often takes several hours to complete, making it impractical for on-site serodiagnosis of infectious diseases in human and wild animals. On-site diagnosis requires a portable immunoassay system that is rapid, sensitive and affordable [2,3], which is also in great demand by medical workers, diagnostic companies, etc. [4]. The advent of microfluidics may provide a possible solution. Various microfluidic mechanisms have been developed to manipulate bio-samples on chip for diagnosis, for instance, mechanical forces, magnetic forces or optical forces. Electric forces are particularly well suited for miniaturization, because a high field strength can be easily generated with low voltages along with microelectrodes and microchannels. This work presents the development of an AC electrokinetics (ACEK) based immunoassay lab-on-a-chip. Through the application of low AC voltages (several volts), AC electrothermal (ACET) effect can be induced to transport

∗ Corresponding author at: 420 Ferris Hall, 1508 Middle Drive, Knoxville, TN 37996, United States. Tel.: +1 8659745494; fax: +1 8659745483. E-mail addresses: [email protected] (X. Liu), [email protected] (K. Yang), [email protected] (A. Wadhwa), [email protected] (S. Eda), [email protected] (S. Li), [email protected] (J. Wu). 1 Both of these authors contributed equally to this work. 0924-4247/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.sna.2011.08.007

and mix the biofluids, which enhances binding due to its swirl-like flow patterns. Our work indicates a ten times acceleration, i.e. the incubation time is shortened from 30 min to around 3 min. AC electro-thermal (ACET) effect can operate even in a highly conductive solution, for which other electrokinetic methods, such as dielectrophoresis, electrophoresis or electroosmosis, do not function effectively [5] [6]. An important benefit from ACET micropumping is the increased binding rate at the channel surface where the electrodes are located. Because ACET microflows will generate swirls towards the electrode surfaces, it helps to break the molecular diffusion barrier that is often encountered in the analysis of diluted samples. Consequently, a shorter incubation time is needed before sufficient binding occurs for reliable detection. Besides ACEK, other strategies were also investigated to provide sample preconcentration in liquids, including field-amplified sample stacking, isotachophoresis, chromatographic preconcentration, and membrane preconcentration. Conventionally, DC electrophoresis is used for charged biomolecule species, and generally such techniques require nanoporous ion-permeable membranes for the operation [7]. Other concentrators utilize free flow electrophoresis, such as electric field gradient focusing [8], field-amplified sample stacking [9], and isotachophoresis [10], capitalizing on spatial gradients of electrophoretic velocity of sample analytes as effected by gradients in ion density, mobility, or solvent viscosity. Many of them require various operational and material constraints, such as special buffer arrangements, reagents, or both. As such, the above preconcentration techniques are difficult

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Fig. 2. Immunoassay protocol. (a) Surface coated with pathogen-specific antigen. (b) Binding of serum antibodies to the antigen. 30 min incubation by pressure driven flow. (c) Wash, then flow in the label (secondary) antibody to bind with the serum antibody. 30 min incubation. (d) Wash, and bound serum antibody was quantified by measuring fluorescent signal from the secondary antibody.

Fig. 1. (a) Schematic of the immunoassay system, and the assembly drawings of (b) light-free box, (c) the holder for the microfluidic chip and detection system and (d) the excitation module.

to implement for on-site and automated applications. Here ACEK methods have minimal requirements on pretreatment of the sample. Additionally, the sensitivity enhancement is independent of the specific type of sensors used, since the enhancement is obtained by sample concentration. Our prior work has performed detailed studies of the ACET effect for in situ concentration of nanoparticles [11] [12], and this work focused on the integration of a miniaturized immunoassay detection system. As schematically shown in Fig. 1, the immunoassay chip electrically controls processes such as reagents delivery, incubations and washing which enables the miniaturization and automation of an on-site immunoassay system. Reagents are preloaded into the reservoirs and are pumped through the reaction chamber sequentially into the waste chamber. Since cross-contamination is a serious concern in biochemical analytical systems, the immunoreactor of our system is designed to be a disposable cartridge. It consists of a poly (dimethylsiloxane) (PDMS) microchannels aligned directly and sealed over an ACET electrodes chip. ACET electrodes are thin film interdigitated electrodes that can be deposited and patterned on silicon, glass or other flat substrates with one photomask process. The PDMS microchannels are formed by soft lithography. So the cartridge can be batch produced at low cost with clean-room technology. The immunoassay described in this work follows the protocol shown in Fig. 2. First, the bottom of the reaction channel is coated with pathogen-specific antigen. Then serum sample is introduced. If the serum sample is obtained from an animal infected with the pathogen, the sample may contain antibodies that bind to

the antigen. On the other hand, serum of an uninfected animal is unlikely to contain such antibodies. Therefore, infection of animal with the pathogen can be diagnosed by quantifying the antibodies bound to the antigens. Excessive (unbound) serum antibodies will be washed away, then fluorescently labeled secondary antibody will be introduced. The secondary antibody binds specifically to the serum antibody. If there is no bound serum antibody, the secondary antibody will be flushed away during the subsequent wash steps. Finally, the bound serum antibody will be quantified by measuring fluorescent signal emitted from the secondary antibody. In a laboratory, fluorescent detection can be accomplished using laser induced fluorescence in coordination with a fluorescence microscope or a photomultiplier, etc. However, such equipment is too cumbersome or expensive to be practical for field testing. In this work, the detection is successfully achieved using a light emitting diode induced fluorescence (LED-IF) with a low cost minispectrometer (RED TIDE USB650, Ocean Optics Inc., Dunedin, FL), in coordination with the ACET lab-on-a-chip. There are previous reports that used LED-IF detection for microfluidic chips [13–17], but the focus of their work is on the set up of the optoelectronic system [18]. This work deals with the operation of the system, including the design consideration for the microfluidic chip and the construction and integration of optoelectronics, so as to realize a portable and rapid immunoassay system. 2. On chip pumping by AC electrothermal effect Electrokinetic pumping is a promising technology for on chip continuous micropumping as it is easy to fabricate, free of moving parts, highly scalable, unsusceptible to particles in the fluid (e.g. cells) and compatible with microchannel integration. However, electrokinetics faces other challenges. A DC pump has severe drawbacks [19], such as the requirements of high voltage (in the kV range, —difficult to integrate into a microsystem), electrode reactions and dissolution (—possible contamination), and being limited to low ionic strength fluids (—so it excludes many biofluid applications). The newly emerged ACEK methods can avoid the above drawbacks of DC pumps because they use high frequency electric fields over an array of interdigitated electrodes to suppress reactions. ACEK includes AC electroosmosis (ACEO) and ACET effect. ACEO cannot be used for biofluid pumping due to the compression of electrical double layer at high ionic concentration in biofluids. It is observed that peak ACEO velocity decreases with increasing conductivity, and ACEO flows are unable to propel fluids with conductivity above 85 mS/m [20] (equivalent to 0.05× phosphate buffered saline (PBS) solution).

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chips. More importantly, the experiments have demonstrated a 10 time reduction in detection time by ACET pumping. 3. Experiments 3.1. Chip design and fabrication

Fig. 3. (a) Simulation of pumping and convection by ACET using Comsol® Multiphysics. (b) Electrodes array.

ACET occurs when inhomogeneous electric fields pass through the fluids. Due to Joule heating, temperature distribution in the fluid will be non-uniform. Therefore, there will be temperature gradient T, which leads to the gradients of fluid conductivity and permittivity. To retain charge neutrality, mobile charges are induced in the fluid and move under the influence of the electric field, leading to fluid motion, i.e. ACET microflows. Therefore, ACET effect is the only electrokinetic mechanism that can be used to manipulate fluids at biologically relevant ionic strengths. It has been reported by Wu group [21,22] as well as by other groups [23,24] that ACET mechanism can be used to handle fluids with conductivity above 0.7 S/m. Though ACET effect is based on Joule heating in biofluids, the temperature rise within the fluid is not necessarily high. It is partly due to fast heat conduction from the channel to the ambience (high surface/volume ratio) and partly due to the fact that heating is highly localized to the regions between the electrodes. ACET flow depends on the thermal gradient not the temperature rise, so localized heating within an electrode gap of several microns will be advantageous in generation ACET flow. The temperature rise in an ACET device is typically around 2–5 K with T around 0.1–1 K/␮m [22–25]. In a lab-on-a-chip, ACET effect can perform two functions, to deliver the sample solution through the microchannel and to actively guide the analyte molecules towards the sensors for in situ concentration so as to achieve rapid detection. One unique feature of ACET microflows is the counter-rotating vortices induced above the electrodes, as shown in Fig. 3, which can provide a transversal velocity across the channel for the analytes to reach the sensors placed between the electrodes. Numerical simulation showed that at 20 ␮m channel height, ACET effect can generate strong vortices against the pressure driven flow and increase the binding rate at the sensor [20]. In our system, ACET pumping is adopted for the disposable immunoassay cartridge, which eliminates the obstacle of interfacing macro-world with microchannels if external pumps are used. The electric connection between the cartridge and the system main body can be achieved with spring-loaded contacts, easy to switch

The immunoassay microfluidic chip consists of PDMS channels sealed over a silicon wafer as shown in Fig. 4(a). Soft lithography was adopted to fabricate the microchannel. The silicon wafer, like Fig. 4(b) forms the channel bottom and the layout of the microchannel is shown in Fig. 4(c). On the left side, three microchambers are designed to house serum, secondary antibody and wash solution, respectively. The serum and secondary antibody will be delivered by ACET electrodes at the bottoms of the dispensing channels into the reaction channel where the binding occurs. The antigens are coated only at the bottom of the reaction channel. All other surfaces are coated with a blocking solution. The channel bottom was covered with interdigitated planar gold electrode arrays as illustrated in Fig. 3(b). The electrodes were deposited by electron beam evaporation, consisting of 100 nm thick Au over 5 nm Ti adhesion layer on oxidized silicon wafer. They were patterned by lift-off process to be 15 ␮m and 5 ␮m wide with 10 ␮m and 30 ␮m separations between them. When the channel is filled with reagents, the input impedance of the ACET pump is measured to be around 580  at the operating frequency of 100 kHz (4294A Precision Impedance Analyzer, Agilent Technologies, Inc., Santa Clara, CA). The ACET electrode array within each channel has separate electrical control. Thus, each dispensing channel can be activated individually to achieve sequential reagent delivery. The reagents are expected to flow from their respective reservoir through the reaction channel then into the waste chamber. To avoid the solution going into other dispensing channels and contaminating the other reagents, different widths and lengths were carefully chosen for the channels so that the difference in the hydrodynamic resistance of the channels can be used for passive valving. The reaction channel is designed to be 2 mm long × 2 mm wide for easy alignment of the optical devices. The dispensing channels are designed to be 20 mm long and 0.5 mm wide. All the channels have the same height of 30 ␮m. The hydrodynamic resistance could be calculated according to the following equation,  = 12L/Wh3ˆ (1–0.630 h/W) [26] where  is fluid viscosity, and L, W and h are the channel length, width and height, respectively. As a result, the dispensing channels pose 41.2 times higher fluid resistance than the reaction channel, so that most of the flow will move through the reaction channel instead of other channels. The length of the channel is much larger than the other dimensions, so the entry and exit effects are insignificant. 3.2. Reagents preparation Reagents: In this study, we used the causative agent and serum samples of Johne’s disease for experiment with ACET-based immunoassay system. Johne’s disease is an economically important infectious disease in dairy industry and is caused by infection of cattle with bacterial pathogen, Mycobacterium avium subsp. paratuberculosis (MAP). Pathogen-specific antigen: MAP was obtained from the USDA (Ames, IA) and was cultured in Middlebrook 7H9 medium (Becton Dickinson Microbiology Systems, Franklin Lakes, NJ) with 10% OADC (oleic acid–albumin–dextrose–NaCl) (Becton Dickinson Microbiology Systems, Franklin Lakes, NJ). The medium was

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Fig. 4. (a) The PDMS microfluidic channels bonded over a silicon substrate to form the immunoassay lab-chip. (b) A 4 Silicon wafer with 8 sets of ACET electrodes for immunoassay lab-chips. The layout of the microfluidic chip is shown as (c), three reservoirs on the left were used for reagents.

supplemented with 2 ␮g/ml of Mycobactin J (Allied Monitor, Fayette, MO). The cultures were maintained at 37 ◦ C without shaking until they reached an optical density of approximately 0.7 at 600 nm. MAP was harvested from the liquid culture at stationary phase and centrifuged at 2600 × g for 10 min; the pellet was resuspended in 80% ethanol, agitated by vortex at room temperature for 2 min, and centrifuged at 10,621 × g for 10 min. The supernatant was collected, diluted in 100% ethanol (1:80) and used as antigen for immunoassay. Serum samples: Bovine serum samples were obtained from the National Animal Disease Center, USDA, Ames, IA. The positive serum was obtained from a dairy cow (breed, Holstein; age, 2.8 years old; gender, female) that was tested positive for Johne’s disease by both bacterial culture method and enzymelinked immunosorbent assay. The negative sample was obtained from a dairy cow (breed, Holstein; age, 2.4 years old; gender, female) that was tested negative for Johne’s disease also by the two procedures above. The aforementioned tests were conducted in the veterinary hospital, the University of Tennessee, prior to the lab-on-a-chip immunoassay. Secondary antibody: The secondary antibody, goat anti-bovine IgG (heavy plus light chain) labeled with the DyLight 488, was obtained from Jackson ImmunoResearch Laboratories (West Grove, PA). Before using in the immunoassay, the secondary antibody was diluted in PBST (10 mM phosphate buffered saline, pH 7.0, containing 0.05 (v/v)% Tween 20).

3.3. ACET-driven immunoassay Before immunoassay test on microfluidic chip platform, the following pre-conditioning steps were conducted. First, 5 ␮L of antigen solution was dropped on top of the electrode at the reaction channel and coated on the channel bottom through passive adsorption. After incubation at room temperature (20–25 ◦ C) for 15 min, the PDMS channel top is sealed over the electrode chip and an initial channel blocking step is performed as follows to reduce unwanted surface absorption of antibodies. Buffer B (10 mM phosphate buffered saline, pH 7.0, containing 0.05 (v/v)% Tween 20 and 10 (v/v)% SuperBlock, PIERCE Biotechnology, Rockford, IL) was loaded in and incubated for 15 min at room temperature. After this pre-conditioning of microchannels, the immunoassay was carried out as described below. The immunoassay was carried out in six steps, including dispensing serum, incubation, flushing unbound serum with PBST, dispensing secondary antibody, incubation and flashing out unbound secondary antibodies with PBST. After loading 5 ␮L of (either positive or negative) serum sample into the reservoir, the solution would be delivered into the reaction/detection channel. An AC signal of 10 Vpp at 100 kHz (Agilent 33220A 20 MHz Function/Arbitrary Waveform Generator, Agilent Technologies, Inc., Santa Clara, CA) was applied for pumping and incubation. As the impedance is not matched to the output impedance of the waveform generator, the actual applied voltage on the electrodes is

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Fig. 5. (a) The testing setup under a fluorescence microscope to observe and evaluate the antigen–antibody binding process. Fig. 3(b)–(c) shows fluorescence images of immunoassay after 3 min of incubation (b) with positive primary antibody and ACET pumping, (c) with negative antibody and ACET pumping, and (d) with positive pumping but pressure driven pumping. Left images of (b)–(d) are photos taken by a fluorescent microscope, and right images of them are the 3D color rendering of the fluorescent intensity distribution over the electrode chip surface.

monitored by an oscilloscope (Tektronix TDS2014, Tektronix Inc., Beaverton, OR) during the experiments. Afterwards, the channel was flushed by PBST twice. Then, 5 ␮L of secondary antibody was loaded into the channel and incubated for 3 min. The channel was again flushed with PBST for 3 times. The total operation takes about 20 min. The detector was then turned on in order to acquire the fluorescent information.

3.4. Optical detection setup Fluorescence detection is the most commonly used method of detection for microfluidic immunosensor systems [2]. DyLight 488, with excitation at 493 nm and emission 518 nm, was conjugated with the secondary antibody as a fluorescence label. The optical path arrangement employed in this study is shown in Fig. 1, with

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a spectrometer as the detector. Fig. 1(c) shows that the excitation module includes an LED with 490 nm wavelength (Aqua 15◦ 5 mm LED, LED Supply, Randolph, VT) which was mounted onto a LED mount (S1LEDM, Thorlabs, Newton, NJ), a shortpass filter with cutoff wavelength at 500 nm (FES0500, Thorlabs, Newton, NJ) and two aspheric lenses for collimating and focusing. The LED was set to operate under ∼25 mA DC current and powered through the USB port of a laptop. The LED was only turned on during the data acquisition to avoid photobleaching of the fluorescence label. After filtered by the shortpass filter, the excitation light beam passes through an aspheric lens (AL1210-A, Thorlabs, Newton, NJ) and an aspherical condenser lens (ACL1512-A, Thorlabs, Newton, NJ). All of these lenses were mounted in a slotted lens tube (SM1L30C, Thorlabs, Newton, NJ) for isolation from ambient light and easy handling, as shown in Fig. 1(d). The lens tube was then positioned at ∼60◦ angle to the plane of the microchannel. The fluorescence emitted from the reagents was collected by an optical fiber at a 45◦ angle to the channel plane. The fluorescence signal goes through a 520 nm long pass film filter (Yellow 12 Kodak Wratten Color Filter, Edmund Optics Inc., Barrington, NJ), and is coupled into the mini-spectrometer. The integration time of the mini-spectrometer was set to 4 s. Then the acquired spectra data were processed by a custom-coded program using Visual Studio 2008. The lateral position of the microfluidic chip can be adjusted so that the LED light spot will be focused over the 2 mm × 2 mm reaction/detection channel. The whole optical detection system is housed in a portable light-free box with a volume of 15 cm × 10 cm × 10 cm, as shown in Fig. 1(b).

4. Results and discussion Two sets of experiments were performed. One was to establish the feasibility of immunoassay using an ACET lab-on-a-chip. Due to the use of electric field, there has been concern of any adverse effect on binding of antigen and antibody or other artifacts. A fluorescence microscope (NIKON eclipse lv-100, Japan) equipped with a CCD camera (Photometrics CoolSNAP ES, Roper Scientific, Germany) and a fluorescence illumination system (X-Cite® 120, EXFO Photonic Solutions Inc., Canada) was used to observe and evaluate the antigen–antibody binding process as shown in Fig. 5(a). Fig. 5(b) and (c) shows the fluorescence images for the positive and negative primary antibodies, respectively, after 3 min of incubation. It demonstrates that the immunoassay procedure is compatible with AC electric field and an ACET based immunoassay chip can differentiate the positive and negative serum samples. Higher fluorescence intensity was observed at the electrode edges. This pattern agrees with the theoretical prediction and numerical simulation. Because electric fields were stronger at the electrode edges, so higher temperature rise/thermal gradients and hence higher flow velocities were expected to occur at the region, and consequently more biomolecules were carried through those regions by microflows for binding reactions. Furthermore, the experiments have proven that ACET lab-on-a-chip could speed up the binding process to achieve accelerated detection. Here, detection was achieved after 3 min of incubation as opposed to 30 min of incubation if pressure driven flow is used. Immunoassay experiments with pressure driven flow was also conducted. As a comparison, Fig. 5(c) shows the case when no electric field was applied. At the end of 3 min, hardly discernable fluorescence can be detected. Because the binding is dependent on the molecular diffusion, the light intensity over the channel bottom is uniform, which is in sharp contrast to ACET based immunoassay. The next set of experiments is detection by LED-IF after 3 min incubation by ACET effect along with control experiments using pressure driven flow. Fig. 6(a) shows the spectral readout from 8

Fig. 6. The spectra (a) collected from reaction and detection channel filled with primary antibody, secondary antibody for 4 samples, (b) background noises deducted and the mean value of positive sample and negative sample. The P# and N# represents the positive and negative samples used by ACET pump while the P (Pressure) and N (Pressure) stands for positive and negative samples driven by pressure.

different experiments with 6 samples pumped by ACET effect (3 positive sera and 3 negative sera) and 2 samples by pressure driven flow (P3 and N3. The spectral response from the negative serum is due to non-specific binding, corresponding to Fig. 5(b), which is almost inevitable during immunoassay. The source of background light could be thermal noise from the optoelectronics, and ambient light leaking into the box through the slot where the immunoassay cartridge is inserted. The background light can be subtracted from the spectral data obtained after an immunoassay to yield smoother curves, as shown in Fig. 6(b). It can be seen that there is substantial difference in the intensity of fluorescence emission between the positive and negative samples by ACET driven, while there is no discernable difference between the samples by pressure driven flow for the same duration. On average, the positive samples were about three times as bright as the negative ones. While a spectrometer is used, it can be readily replaced by a photodiode. The output from a photodiode after a bandpass or long pass filter will correspond to the area under the intensity curve defined by the filter. For the dye used here, the filter will have cut-off wavelength at 515 nm. The spectral responses in Fig. 6 exhibit sufficient intensity differentiation at wavelength longer than 515 nm to warrant detection by fluorescence intensity. This experiment serves as the proof of principle that serodiagnosis of Johne’s disease can be achieved using

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LED-IF within 3 min’ incubation for each step. Further development includes actual implementation with a photodiode. The optical LOD (limit of detection) were measured with fluorescent microbeads (FluoSpheres® carboxylate-modified microspheres, 0.2 ␮m, yellow-green fluorescent (505/515) *2% solids*, Molecular Probes, Invitrogen, US) using the same setup for immunoassay. The LOD is determined to be ∼0.5 nM. The bovine serum (primary antibody) contains 10–20 mg/ml of IgG (∼1 mmol/l), far exceeding the LOD of our system, so this lab-on-achip system is adequate for immunoassay. The developed lab on a chip immunoassay system will have a small form factor. The only benchtop instrument used here is the waveform generator, which can be easily replaced by an off-theshelf device, for example, USBwave12 by Elan Digital Systems Ltd., UK. Integrating all the components and providing automated data acquisition and processing will be the next stage of development. 5. Conclusion This paper presents the development and experiments of an immunoassay lab-on-a-chip for rapid detection of antibody binding to pathogen-specific antigen, suggesting that this system may be used for serodiagnosis of infectious diseases. We have demonstrated proofs of concept for all the critical technologies involved. We have experimentally differentiated the positive and negative samples as shown in Figs. 5 and 6 with the following improvements, (1) using an LED as excitation source (costs: ∼$1 versus >$2,000 for conventional lamp); (2) detection acceleration to within 2–5 min of incubation versus traditional 30 min (twice); and (3) the development of a low cost, disposable, and ultra portable immunoreaction chip. The fluorescence immunoassay developed in this study showed high analytical sensitivity and was much faster than conventional methods. To our knowledge, this work is the first to use active guidance of analytes diffusion to nanoscale sensors, which previously have been purely based on thermodynamic diffusion in a flow through channel. Acknowledgements This study was supported by the U.S. National Science Foundation under Grant No. ECS-0448896, the Research Fund for the Doctoral Program of Higher Education of China under Grant No. 20090191110026, and M-CERV program and UTRF Maturation Fund of The University of Tennessee. Microfabrication of this research was conducted at the Center for Nanophase Materials Sciences, which is sponsored at Oak Ridge National Laboratory by the Scientific User Facilities Division, U.S. Department of Energy. X. Liu and S. Li gratefully acknowledge the financial support from China Scholarship Council. References [1] Q. Xiang, G. Hu, Y. Gao, D. Li, Miniaturized immunoassay microfluidic system with electrokinetic control, Biosensors and Bioelectronics 21 (10) (2005) 2006–2009. [2] A. Bange, H.B. Halsall, W.R. Heineman, Microfluidic immunosensor systems, Biosensors and Bioelectronics 20 (12) (2005) 2488–2503. [3] W.G. Lee, Y.G. Kim, B.G. Chung, U. Demirci, A. Khademhosseini, Nano/microfluidics for diagnosis of infectious diseases in developing countries, Advanced Drug Delivery Reviews 62 (4–5) (2010) 449–457. [4] C.C. Lin, J.H. Wang, H.W. Wu, G.B. Lee, Microfluidic immunoassays, Journal of the Association for Laboratory Automation 15 (3) (2010) 153–274. [5] A. Ramos, N.G. Green, A. Castellanos, Fluid flow induced by nonuniform ac electric fields in electrolytes on microelectrodes. II. A linear double-layer analysis, Physical Review 61 (4) (2000) 4019–4028. [6] J. Wu, AC electro-osmotic micropump by asymmetric electrode polarization, Journal of Applied Physics 103 (2) (2008) 024907-5.

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Biographies Xiaozhu Liu is a Ph.D. candidate of OptoElectrical Engineering at Chongqing University (CQU), and a visiting student in the University of Tennessee, Knoxville (UTK). His research interests are in microfluidics, optoelectronics and their applications. Kai Yang received his Ph.D. in Electrical Engineering from The University of Tennessee, Knoxville (UTK), USA, 2011. He received his Bachelor degree from Shanghai Jiao Tong University (SJTU), Shanghai, China, and Master of Science degree from UTK, both in Electrical Engineering. His research interests are in micro/nano-fluidics with AC electrokinetics, biosensor and immunosensor design for lab-on-a-chip applications. Ashutosh Wadhwa is a Ph.D. candidate associated with the Center for Wildlife Health, University of Tennessee Knoxville. He received his Bachelor’s degree in Veterinary Science and Animal Husbandry (B.V.Sc & A.H.) from the College of Veterinary Science and Animal Husbandry, Mathura, India and Masters in Epidemiology and Preventive Veterinary Medicine at the College of Veterinary and Animal Science, Bikaner, India. His research interests are epidemiology and developing better diagnostic test for animal and human infectious diseases. Shigetoshi Eda is an Associate Professor in the Department of Forestry, Wildlife and Fisheries at the University of Tennessee Knoxville. He received his Ph.D. degree from the Tokyo University of Pharmacy and Life Sciences in 1997. His research interest is development of immunoassays for diagnosis of infectious diseases in animals and humans.

X. Liu et al. / Sensors and Actuators A 171 (2011) 406–413 Shanshan Li is a Ph.D. candidate of Mechanical Engineering at Harbin Institute of Technology (HIT), and a visiting student in the University of Tennessee, Knoxville (UTK). Her research interests are in microfluidics and microparticle manipulation. Jie Wu received the Ph.D. degree in Applied Physics from the Chinese Academy of Sciences, China, in 1999, and the Ph.D. degree in Electrical Engineering from the University of Notre Dame, Notre Dame, IN, in 2004. From August 2003 to July 2004,

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Dr. Wu was a Postdoctoral Research Fellow with the Center for Microfluidics and Medical Diagnostics, Department of Chemical and Bio-molecular engineering, University of Notre Dame. Currently, she is an Associate Professor with the Department of Electrical Engineering and Computer Science, University of Tennessee, Knoxville. Her research interests are solid-state device physics, bio micro electromechanical systems, microfluidic sensors and actuators.