PAPER
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Self-referencing a single waveguide grating sensor in a micron-sized deep flow chamber for label-free biomolecular binding assays Po Ki Yuen,* Norman H. Fontaine, Mark A. Quesada, Prantik Mazumder, Richard Bergman and Eric J. Mozdy Received 25th January 2005, Accepted 28th June 2005 First published as an Advance Article on the web 21st July 2005 DOI: 10.1039/b501219h In order to allow the design of increasingly sensitive label-free biosensors, compensation of environmental fluctuations is emerging as the dominant hurdle. The system and technique presented here utilize a unique combination of microfluidics, optical instrumentation, and image processing to provide a reference signal for each label-free biomolecular binding assay. Moreover, this reference signal is generated from the same sensor used to detect the biomolecular binding events. In this manner, the reference signal and the binding signal share nearly all common-mode noise sources (temperature, pressure, vibration, etc.) and their subtraction leaves the purest binding signal possible. Computational fluid dynamic simulations have been used to validate the flow behavior and thermal characteristics of the fluids inside the sensing region. This system has been demonstrated in simple bulk refractive index tests, as well as small molecule (biotin/streptavidin) binding experiments. The ability to perform not only simple binding but also control experiments has been discussed, indicating the wide applicability of the technique.
Introduction While fluorescent and radioactive detection technologies still dominate drug discovery and biological research worldwide, an increasing demand has emerged for label-free technologies that avoid typical fluorescent or radioactive assay detriments, such as the potential to alter fundamental molecular function, complicated waste disposal or significant time and effort in assay development.1 Even though label-free technology has proven its ability to quantify molecular binding, the drive toward increasingly sensitive detection (e.g., smaller molecular weight binders) has exposed some significant hurdles to next-generation systems.2 In particular, when label-free sensors are designed to be more sensitive to biomolecular species, they likewise display heightened sensitivity to unwanted environmental perturbations.3,4 As a result, ambient pressure, temperature, etc. changes can present more signal than the actual biomolecular binding event. In such a situation, there may not be any advantages to further optimize the sensor itself; but one must rather deal with mitigation of environmental effects. In the case of optical waveguide sensors, the primary noise and drift sources are temperature fluctuation, mechanical vibration, thermal expansion, and laser source drift.3,4 An additional source of drift can include the presence of microporosity in the waveguide sensor.5 One means of dealing with environmental drift is to use temperature control, like that in existing instruments (e.g., Biacore AB, Uppsala, Sweden). However, instruments with active thermal management are expensive; moreover, Science and Technology, Corning Incorporated, Corning, New York, 14831-0001, USA. E-mail:
[email protected]; Fax: (607) 974-5957; Tel: (607) 974-9680
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temperature control alone cannot correct for all environmental factors. Alternatively, one can diminish the impact of environmental fluctuations by providing a referencing method using more than one sensor where the difference between sensor signals is the measurable, insensitive to common-mode noise. Goddard et al. have demonstrated that, by incorporating an additional buried resonant mirror waveguide layer into a surface resonant mirror sensor structure, an internally referenced resonant mirror sensor can be generated.3,4 However, by using more than one sensor, the number of processing steps and/or the required detection optics may increase with the number of sensors added. As a result, there may be a physical limitation on the number of sensors that can be added and/or how close the sensors can be positioned together in a device. Also, physically separated sensors may experience different environmental fluctuations and may have different characteristics and performances. All of these factors add to the uncertainty and hence can limit the accuracy of the referencing technique. In addition to the need for environmental insensitivity, referencing is also often required in control experiments used to uncover false positives, to eliminate the effect of the bulk fluid refractive index, or to perform signal subtraction in biomolecular binding assays. An example of such a control experiment is the determination of whether specific or nonspecific binding of biomolecules occurred in a surface-binding assay; this requires two or more sensing regions to allow for different surface chemistries, target analytes, or probes. In other assays, one may purposely prepare two sensors to provide different, yet predictable responses, or perhaps use the separate sensors just to eliminate bulk refractive index effects. In all of these control experiments, it is again advantageous to have reference sensing regions as physically close as possible to mitigate unwanted environmental effects as well as allow Lab Chip, 2005, 5, 959–965 | 959
precise control over combined reaction environments. Since sensor performance may vary due to fabrication artifacts, the use of multiple sensors distributed over a large area may yield less than optimal referencing results. Accordingly, there is a need for a technique to address the aforementioned shortcomings and other shortcomings of traditional label-free instruments. The label-free system described in this paper, based upon grating-coupled optical waveguide sensors,6 satisfies these needs and furthermore enables novel control assays by providing self-referencing with a single sensor. The single waveguide sensor is effectively divided into two halves by microfluidic fluid flow, and the optical signals from each half of the sensor form the reference pair. This is made possible through careful design of the micron-sized channel, as well as a clever use of optical instrumentation and software to provide an efficient use of information from a single sensor. By using one sensor to generate both detection and reference signals, the fabrication or environmental differences commonly found in typical referencing schemes are avoided.
Materials and methods Microfluidic device design The H-shaped microfluidic device consists of a single 100– 200 mm deep multiple flow channel located over a 3 mm 6 3 mm waveguide grating sensor that is used for detecting biomolecular binding events (Fig. 1). The microfluidic channel has two inlets and two outlets, allowing several configurations of fluid flow over the sensor. Two fluids can be flowed simultaneously into the sensing region by opening the two inlets (Fig. 2A). Also, at any given time, one of the inlets can be closed such that only one fluid can flow through the sensing region (Fig. 2B). Microfluidic device fabrication Soft lithography7 was used to fabricate the H-shaped microfluidic device. A photo mask of the device was generated from a CAD file. After spin coating a 100–200 mm thick SU-8 photoresist (MicroChem Corp., Newton, MA) on a glass wafer, the mask was used in 1 : 1 contact photolithography with the SU-8 photoresist to generate a negative ‘‘master’’ template of the device, consisting of patterned photoresist on
Fig. 2 Computational fluid dynamic simulations of the particulate flow path inside the H-shaped microfluidic device. The device was 100 mm deep, and held at 27 uC. (A): Inlet 1: water flow rate was 485 mL min21 at 26 uC. Inlet 2: water flow rate was 485 mL min21 at 28 uC. (B): Inlet 1 was closed. Inlet 2: water flow rate was 485 mL min21 at 28 uC.
the glass wafer. Positive replicas, 1 mm thick, with embossed structures were fabricated by casting and curing (e.g., at 65 uC for one hour) an optically transparent prepolymer of silicone or polydimethylsiloxane (PDMS) (e.g., Sylgard1 184 Silicone Elastomer from Dow Corning Corporation, Midland, MI) against the master template. After releasing the replica from the master, the replica was cut to the right dimensions using a surgical blade. Inlets and outlets were punched into the replica using a sharpened 2 mm diameter hollow tube. Finally, to complete the fabrication process, the processed replica was laid on the waveguide grating sensor to close the open channels. Waveguide grating sensor design and fabrication
Fig. 1 Schematic of the H-shaped microfluidic device.
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A typical grating-coupled waveguide sensor consists of an optical waveguide fabricated on a transparent dielectric substrate, and includes a diffraction grating fabricated above, below, or even within the waveguide. This grating serves to couple the waveguide mode to an external light source and detector, allowing the remote monitoring of the waveguide mode. The dimensions of the waveguide grating sensor used in this paper were 3 mm (width) 6 3 mm (length) 6 150 nm (depth). The grating grooves were rectangular with a period of 500 nm, a depth of 50 nm, and a 50% duty cycle. A conventional UV cast and cure process was used to fabricate the grating structure.8 In the cast and cure process, a negative This journal is ß The Royal Society of Chemistry 2005
master stamping tool, consisting of patterned grating structures, must first be fabricated using a technique such as E-beam lithography. A UV curable polymer is then applied to the surface of a glass substrate to obtain a uniform thickness. Next, the master stamping tool is pressed into the polymer to mold a positive replica of the grating structure on the glass substrate. While the stamp and substrate are in contact, the polymer is cured under a UV source, and the tool subsequently released. Finally, a high refractive index thin film is deposited on the sensor to form the waveguide; the specific device used in this paper received a 150 nm thick tantalum oxide (Ta2O5) coating. Optical detection system When the surface of a grating-coupled waveguide sensor is used as the platform for biomolecular binding assays, the small refractive index changes that accompany binding events cause changes in the waveguide mode, subsequently causing a change in the angle or wavelength of the light remotely coupled into/out of the sensor by the grating. By monitoring the coupling wavelength or angle, one may thereby quantify binding at the surface of the sensor. The sensor measurement system included an interrogation system that directs an input optical beam from a laser source onto the waveguide grating sensor and receives the sensor’s output signal (Fig. 3). The monochromatic input optical (laser) beam is focused into a line approximately 4 mm (wide) 6 0.2 mm (long), which spans the entire width of the sensor. The focusing of the beam creates a range of angles to interact with the grating-coupled waveguide sensor, where only a resonant angle will couple with the waveguide. The line illumination is in contrast to more conventional spot illumination approaches,9 and allows the generation of many angles while still covering the width of the sensor. The resonant angle that couples into the waveguide also couples back out of the sensor, forming a resonant reflection response from across the width of the waveguide grating sensor; this resonant angle is identified by its position in the far field, as tracked by a CCD camera. The change in location of the resonance image at the camera indicates an angular response shift that is due to a change in the waveguide’s surface index of refraction: a change in the bulk
Fig. 3 Schematic of the self-referencing optical detection system.
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assay fluid refractive index and/or a change in biomolecular binding/adsorption on the surface. A calibration plot of the sensor response as a function of the bulk fluid refractive index for a grating-coupled waveguide sensor (Fig. 4) illustrates the high linear response of the sensor with an R2 value of 0.9997 (Format Trendline function from Microsoft Excel). Since the response across the sensor’s width maps to location across the CCD camera, software may subsequently be used to parse the CCD image and analyze the separate responses from each half of the sensor. By properly flowing sample and reference fluids across each half of the sensor within the micron-depth channel, detection and reference signals may be associated with each half of the sensor. Given that each half of the sensor is subject to the same environmental conditions while experiencing different assay fluids, taking the difference of the responses from each half of the sensor will result in the referenced binding signal of interest. Furthermore, if one desires, the sensor can be divided into an arbitrary number of distinct regions with a similar combination of fluid flow and software image analysis. Computational fluid dynamic simulation In order to validate the flow behavior and thermal effects of the fluids inside the H-shaped microfluidic device, computational fluid dynamic (CFD) simulations were performed using commercial software (Fluent, Inc., Lebanon, NH). For each simulation, the flow rates and the temperatures of the two flow regions were varied while the temperature of the device (micron-depth channel) was fixed at a constant temperature of 27 uC. Self-referencing testing Three different flow experiments were performed to demonstrate the feasibility of the self-referencing technique. These were bulk refractive index tests, a biotin/streptavidin flow binding experiment, and a second biotin/streptavidin binding experiment where surface immobilization (instead of flow) differentiated the binding/non-binding regions of the sensor.
Fig. 4 Calibration plot of the sensor response as a function of the bulk fluid refractive index for a grating-coupled waveguide sensor, illustrating the high linear response of the sensor with an R2 value of 0.9997.
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The experimental process is described here, and the results are treated in the next section. In the bulk fluid refractive index experiments, fluids were pumped through the H-well and across the sensing region using a two-syringe infusion/withdrawal pump (Model no. SP210iw, World Precision Instruments, Sarasota, FL). The experiment involved first flowing deionized water through only one inlet to flood the entire sensor region, followed by the simultaneous insertion of a glycerol solution flow through the second inlet (to pass over half of the sensor), finished by a return to the solitary deionized water flow as a final rinse. This test served to quantify the sensitivity of the referenced system to pure refractive index changes. Alternatively, the experiment could have involved the continuous use of both inlet ports, where one flow stream of deionized water could have been briefly switched with glycerol solution, thereby maintaining two flow streams inside the sensing region at all times. The first biotin/streptavidin binding experiment (molecular weights of 244 Da and y60 kDa, respectively) was performed to demonstrate the sensitivity of the self-referencing technique. In the sensor preparation phase, the sensor was initially washed with deionized water, isopropyl alcohol (Fisher Scientific International Inc., Hampton, NH), and again with deionized water. Then, 5% aminopropylsilsesquioxane (APS) (Gelest Inc., Morrisville, PA) in water was added to the sensor and incubated for 10 min at room temperature. Next, the sensor was washed several times with deionized water before a final ethyl alcohol (Fisher Scientific International Inc., Hampton, NH) wash, and was immediately blown dry with nitrogen. A 2.5 mg mL21 streptavidin solution (Sigma Chemical Co., St. Louis, MO) was then added to the sensor and incubated for 30 min at room temperature. Next, the sensor was washed several times with 16 phosphate buffer saline (PBS) solution (Sigma Chemical Co., St. Louis, MO). After manually immobilizing streptavidin onto the whole sensor surface as described above, a fluid flow sequence similar to the bulk refractive index test was employed, where a 16 PBS solution was used instead of deionized water, and biotin (in PBS) solution (Sigma Chemical Co., St. Louis, MO) was used instead of the glycerol solution. Because only the biotin would specifically bind to the streptavidin, the 16 PBS solution served to generate a reference signal to the biotin binding signal. The second biotin/streptavidin binding experiment again demonstrated the system’s sensitivity, using an alternative fluid flow format. In the sensor preparation phase, after applying the 5% APS surface chemistry, a continue side-by-side flow method was used to immobilize the separate halves of the sensor surface with streptavidin and bovine serum albumin (BSA) (Sigma Chemical Co., St. Louis, MO) respectively, to create binding and non-binding regions on the sensor surface. BSA effectively blocks the binding of biotin to one half of the sensor surface, thereby creating a reference to the biotin/streptavidin binding signal from the other side of the sensor. Because this surface chemistry treatment distinguishes the two halves of the sensor, a single fluid flow from one inlet port could therefore be employed throughout the course of the experiment. First, a 16 PBS solution was flowed in order to establish the baseline signals, and then biotin solution was 962 | Lab Chip, 2005, 5, 959–965
flowed to create a differential binding signal. In this case, the final 16 PBS solution wash was not necessary since any non-specific binding/bulk refractive index changes would be referenced out by the different surface chemistries. Thus, the referenced signal (the biotin binding signal minus the nonbinding BSA reference signal) indicated only the effect of the biotin binding.
Results and discussion Computational fluid dynamic simulation The primary goal of the fluidic configuration employed in this paper is to remove environmental effects (e.g. temperature) while providing high-quality, versatile assay data. This requires that the two separate fluids flowed over the sensor rapidly thermalize yet do not physically mix. Fortunately, careful fluidic channel design accomplishes this goal, as verified by the CFD simulations (Fig. 2A). The small length scale of the channel precludes any possibility of eddy diffusivity due to turbulence and/or shear layer instability between the two fluids, and the only means of mixing is therefore molecular diffusion. Moreover, the mass diffusivity of the molecular entities of the two fluids typically employed for biomolecular binding assays is many orders of magnitude smaller than the thermal diffusivity of the two fluids. This leads to disparate thermal and mass diffusion length scales in the channel. Due to the low mass diffusivity, the two streams are separated by a very thin diffusion (mixing) layer, which grows relative to the amount of time the two fluids are in pffiffiffiffiffiffiffiffi contact. The width of the diffusion interface is 2Dt where D is the mass diffusion coefficient and t is time (t 5 L/U, where L is the distance of the sensing region from the inlet, and U is the average flow velocity). The small value of D (e.g., for y850 Da fluorescein biotin and y66 kDa bovine serum albumin, D 5 3.4 6 1026 cm2 s21 and 6.5 6 1027 cm2 s21, respectively)10 ensures that the chemical/compositional integrity of each stream will be maintained except in a very thin layer near the flow-stream interface. On the other hand, the relatively large value of thermal diffusivity ensures that the lateral temperature distribution in the channel ispuniform (data not shown): the thermal boundary layer ffiffiffiffiffiffiffi grows as 2at, where a is the thermal diffusivity (e.g., a 5 1.4 6 1023 cm2 s21 for water). Compounding the benefit, the dominant heat transfer occurs between the fluid and the micron-depth channel, where the thermal diffusion length is the height (in microns) of the channel, whereas the mass diffusion occurs between the two fluid flows, and its length extends the entire width (in millimeters) of the channel. By exploiting these disparate diffusion length scales of heat and mass, this self-referencing technique is able to use one sensor to investigate interactions between biomolecules in the sensing area. As Fig. 2A and 2B show, two inlet fluids may flow side-by-side with little mass diffusion, or one inlet of the micron-depth channel may be used alone to fill the entire sensor chamber with a single fluid. These two modes allow the device to accomplish an exchange of fluids over the two halves of a single sensor and thereby accomplish referencing with one sensing device. This type of self-referencing technique effectively reduces or removes the sensitivity that the sensor has to perturbations in angle, location, temperature, source This journal is ß The Royal Society of Chemistry 2005
wavelength, non-specific binding between biomolecules, as well as the effect of the bulk fluid refractive index. Self-referencing testing The feasibility of the self-referencing technique was first demonstrated experimentally with two bulk fluid refractive index experiments (Figs. 5 and 6). When deionized water was initially flowed into the whole sensing region of the sensor (Fig. 5A, t , 23 s), the two baseline signals corresponding to the left and right halves of the sensor had the same response characteristics. Both of them gradually drifted downward with small synchronized fluctuations, demonstrating the impact of environment, and their standard deviations (t 5 0–20 s) are shown in Fig. 5A (calculated with a running root-mean-square algorithm using STDEV function from Microsoft Excel). At t 5 23 s, glycerol solution was introduced into the sensing region through the left inlet (Inlet 1) without interrupting the flow of deionized water at the right inlet (Inlet 2). The signal corresponding to the left half of the sensor (the pink curve) rose rapidly without affecting the signal in the right half of the sensor (the dark blue curve), indicating the higher refractive index of the glycerol solution. For illustrative purposes, the sensor resonance image before and after the introduction of a high refractive index glycerol solution in a similar experiment is depicted in Fig. 6. The shift in the left half of the recorded resonance image is easily observed. This image also confirms that the two fluids were flowing side by side inside the sensing region over each half of the sensor and that a thin diffusion
Fig. 5 Resonance plots during a bulk fluid refractive index experiment. (A): First, deionized water flowed over the whole sensing region of the waveguide grating sensor (refer also to the color-coded inserts in the plot). Next, glycerol solution was introduced into the left half of the sensing region through the left inlet (Inlet 1) without interrupting the flow of deionized water from the right inlet (Inlet 2). Finally, the flow stream of glycerol solution was stopped while maintaining the flow of deionized water (through the entire channel). The pink and dark blue curves correspond to the left and right half of the sensor, respectively. (B) The self-referenced curve was generated by subtracting the deionized water signal (dark blue signal) from the glycerol solution signal (pink signal). The flow rate for this experiment was 350 mL min21.
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Fig. 6 The resonance image of the waveguide grating sensor during a bulk fluid refractive index experiment. (A): Deionized water flowing over the sensor at 20 mL min21. (B): Deionized water and glycerol solution flowing side-by-side over the sensor at 20 mL min21.
layer was captured in the resonance image as a sharp (but not infinite) resonance slope in the middle of the image (Fig. 6B). Returning to Fig. 5A, the flow of glycerol solution was stopped at t 5 50 s and only deionized water was allowed to flow into the whole sensing region. The pink signal dropped rapidly back to the same level as the dark blue signal, indicating that the bulk fluid refractive index returned to the initial value. By subtracting the deionized water signal (dark blue signal, right sensor half) from the glycerol solution signal (pink signal, left sensor half), a self-referenced signal of bulk fluid index change was obtained (Fig. 5B). This self-referenced signal did not exhibit any baseline drift in contrast to the unreferenced signals of Fig. 5A. Furthermore, the standard deviation of the baseline signal fluctuations was reduced by a factor of 2.5 (t 5 0–20 s and t 5 60–80 s). Although the standard deviation of the referenced red signal (Fig. 5B) was the same as the unreferenced pink signal (Fig. 5A) (t 5 30–50 s), it is believed that the use of two separate plastic syringes for the side-by-side flow resulted in a larger uncorrelated fluid pumping noise than found in the singleflow baseline signals. This situation may be remedied by appropriate pump design (e.g., using a mechanical actuator for all flows), such that self referencing can remove the truly common-mode signal. Despite this moderate noise level, this self-referenced signal did not have any of the drift (slope) clearly seen in the unreferenced pink signal (Fig. 5A). The decrease in baseline fluctuation equates to a 2.5 times increase in the label-free system’s index of refraction detection sensitivity, in addition to the aforementioned reduction of environmental drift. Lab Chip, 2005, 5, 959–965 | 963
The self-referencing technique was also demonstrated for small molecule (biotin/streptavidin) binding (Figs. 7 and 8). In Fig. 7, the pink curve corresponds to the raw biotin binding signal from one half of the sensor, while the dark blue signal corresponds to the 16 PBS reference signal from the other sensor half. The referenced biotin binding signal (raw biotin 2 PBS signal) is shown in red. Fig. 7A shows a very good response with a standard deviation of y0.009 before the injection of biotin solution at t , 155 s and after the final 16 PBS wash at t . 580 s. After the final 16 PBS wash at t . 580 s, the net shift of y0.1 response unit for the referenced biotin binding signal (red curve) indicates that binding has occurred. By equally adjusting the baselines (subtracting the upward drift slope) of both the 16 PBS and the biotin raw signals (dark blue and pink curves in Fig. 7B), it can be observed visually that the baseline-adjusted biotin binding Fig. 8 Data obtained from a second biotin/streptavidin binding experiment. The two separate halves of the sensor surface were first immobilized with streptavidin (30 mg mL21) and BSA (30 mg mL21), respectively (refer also to the color-coded inserts in the plot). The BSA created a non-binding reference surface on half of the sensor. A 16 PBS solution was flowed across the entire sensing region of the device, establishing the baseline signals. Then, 140 mM biotin solution was flowed across the entire sensing region, replacing the 16 PBS solution. The pink and the dark blue curves correspond to the raw biotin binding signal and the non-binding BSA reference signal, respectively. The red curve corresponds to the referenced biotin binding signal (pink curve minus dark blue curve). The flow rate was 100 mL min21.
Fig. 7 Resonance plots from the first biotin/streptavidin binding experiment. After manually immobilizing streptavidin (2.5 mg mL21) onto the whole sensor surface, a 16 PBS solution was flowed into the sensing region to establish baseline signals. Next, 140 mM biotin (in PBS) solution and the 16 PBS solution were flowed side by side into the sensing region of the sensor. The 16 PBS solution was used as a reference signal to the biotin binding signal. Finally, the flow stream of the biotin solution was stopped while the 16 PBS solution was continuously flowed across the width of the sensing region to wash off any unbound biotin. The pink and the dark blue curves correspond to the raw biotin binding signal and the 16 PBS reference signal, respectively. The red curve corresponds to the referenced biotin binding signal (pink curve minus dark blue curve). (A): Raw data. (B): The baselines of both the 16 PBS reference (dark blue signal) and raw biotin binding (pink signal) curves in (A) were adjusting equally by subtracting the linear upward drift of both signals. The flow rate was 50 mL min21 for each fluid.
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signal (the pink curve) and the referenced binding signal (the red curve) have a very high degree of correlation (Fig. 7B). This indicates that both the 16 PBS reference signal and the biotin binding signal (Fig. 7A) were affected by the same environmental fluctuations, as a simple linear baseline adjustment was sufficient to correct the signal response. However, in most cases (see Fig. 8) environmental fluctuations will not be purely linear and the self-referencing technique described herein would be the only means to efficiently reference the data. The second biotin/streptavidin binding assay is shown in Fig. 8, where the referenced biotin binding signal (red curve) is the difference between the biotin binding signal (pink curve) and the non-binding BSA reference signal (dark blue curve). The referenced binding signal shows a very good response with a standard deviation (noise) of approximately 0.007 response units. There was a binding shift of approximately 0.1 response units after the biotin injection, commensurate with the shift obtained in the previous binding experiment (Fig. 7A and 7B). Thus, for both experiments, a small but observable shift of approximately 0.1 response units indicates the ability of the label-free sensor system to observe small molecule (biotin/streptavidin) binding. Discussion of potential pitfalls It is expected that the presence of air bubbles would have an adverse effect on the performance of any microfluidic system. In the system described in this paper, any air bubbles formed in the detection area will result in erratic signals due to the This journal is ß The Royal Society of Chemistry 2005
significant difference in refractive indices between air and fluid (air y1.000 versus water y1.333 at room temperature). Also, the presence of air bubbles inside the device may affect the flow pattern of the two adjacent fluids in the micron-depth channel, leading to unbalanced flow profiles and inaccurate self-referencing. The main approaches to preventing bubble formation include the prevention of dead zone within the fluidic path (largely accomplished through channel design), the use of reputable components, and proper loading of fluid samples. Using these techniques, the occurrence of bubbles in the aforementioned experiments was rare, and easily remedied with a quick purge of the system.
Conclusions The system and technique described in this paper utilize a unique combination of microfluidics, optical instrumentation, and image processing to provide a reference signal for each label-free biochemical assay. This technique avoids the expensive and cumbersome requirement for sensitive environmental control in the label-free instrument. Contrary to other referencing schemes employed to date, the reference signal of this work is generated from the same sensor used to detect the biochemical binding events. In this manner, the reference signal and binding signal share nearly all common-mode noise sources (temperature, pressure, vibration, etc.) and their
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subtraction leaves the purest binding signal possible. Computational fluid dynamic simulations have been used to validate the flow behavior and thermal characteristics of the fluids inside the sensing region. This system and technique have been demonstrated in simple bulk refractive index tests, as well as small molecule binding experiments. The ability to perform not only simple binding but also many control experiments has been discussed, indicating the wide applicability of the technique.
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