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A compact and multichannel optical biosensor based on a wavelength interrogated input grating coupler
Sensors and Actuators B 161 (2012) 721–727
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Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb
A compact and multichannel optical biosensor based on a wavelength interrogated input grating coupler Sonia Grego ∗ , Kristin H. Gilchrist, James B. Carlson, Brian R. Stoner Center for Materials and Electronic Technologies, RTI International, 3040 E. Cornwallis Road, Research Triangle Park, NC 27709-2194, USA
a r t i c l e
i n f o
Article history: Received 14 August 2011 Received in revised form 27 October 2011 Accepted 7 November 2011 Available online 25 November 2011 Keywords: Optical waveguide Tunable laser Grating Multichannel sensor Portable biosensor
a b s t r a c t The design, fabrication and characterization of a compact optical biosensor based on an input grating coupler are presented. The sensor device consists of a waveguide-integrated single grating structure. The optical reader is based on the use of a largely tunable, small footprint infrared laser for wavelength interrogation of the sensor and on edge-detection of waveguided modes, resulting in a compact configuration. A multiplexing approach is illustrated and a two-output sensor system is demonstrated. The device is packaged in a microfluidic cartridge for sample handling. The volumetric sensing performance of the system is characterized and a limit of detection of 1 × 10−5 RIU is reported while the surface sensing capability is demonstrated by detection of protein diluted down to 200 ng/ml. © 2011 Elsevier B.V. All rights reserved.
1. Introduction Label-free optical sensing technologies based on detection of molecular binding events have been extensively investigated as a rapid analytical method for biological, medical and environmental applications. These approaches detect biomolecule absorption on the transducer surface as a change in the refractive index at the transducer–liquid interface using evanescent waves in either all dielectric or dielectric-metal devices [1,2]. Multichannel biosensor instruments reading receptor-functionalized chips have been demonstrated as powerful immunosensing tools but are limited to laboratory use due to their large size and complexity. Fielddeployable platforms are highly desirable, given the biosensor capability to quickly detect contaminants in food [3,4] or diagnose infections from a clinical specimen such as blood [5], however, they often provide reduced sensitivity compared to benchtop instruments. In recent years, efforts have been reported to develop portable sensors using fully packaged approaches based on surface plasmons [6–8] or grating couplers [9] or to reduce size by integration of sample handling on-chip (e.g. in ring resonators [10]) yet these systems do not achieve a truly hand-held size or do so at expense of lack of multiplexing or non-scalable device manufacturing [11,12]. We present a grating coupler sensor based on wavelength interrogation and edge detection which allows construction of a
∗ Corresponding author. Tel.: +1 919 248 4181; fax: +1 919 248 1955. E-mail address: [email protected] (S. Grego). 0925-4005/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.snb.2011.11.020
compact and rugged sensor suitable for field use. The approach is referred to as Tunable Wavelength Interrogated Sensor Technology (TWIST) and relies on the use of a largely tunable yet compact infrared wavelength source. We previously fabricated gratingintegrated waveguide devices and characterized their sensing performance at = 1.5 m using benchtop components including a wavelength swept laser source for telecommunication testing [13]. The advantage of a broadband wavelength sweeping source for optical waveguide grating sensors was also demonstrated by others [14] by using a custom-assembled fiber-based laser source implemented with an electronically tunable Fabry–Perot filter and other components. That work demonstrated a large wavelength scan at kHz speed driven by an external voltage source and was realized as a benchtop sensor apparatus. This paper describes a miniaturized system serving as optical reader of the waveguide devices including a new optical source and photodetector as well as multichannel capability. The configuration is designed and the components are selected with the ultimate objective of integration into a hand-held device interrogating disposable sensor chips. Major requisites of the system include the use of compact and inexpensive components and the ability to multiplex sensors on the same device to enable on-chip referencing. The waveguide grating sensor chip was integrated with a fluidic cartridge and the system detection capabilities were characterized in terms of volume refractive index sensing and surface sensing of IgG antibody. The miniaturized system performance is improved compared to our previously reported values to a detection limit of 1 × 10−5 RIU and an on-chip reference configuration is demonstrated.
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Fig. 1. Schematic of the TWIST sensor principle.
2. Tunable wavelength interrogated sensor technology (TWIST) 2.1. Principle of operation TWIST is based on evanescent wave sensing of surface binding analytes using a wavelength-tunable, input grating-coupler configuration. A schematic of the sensing principle is shown in Fig. 1. A collimated laser beam is incident on the grating, coupled into the silicon oxynitride waveguide film, and emerges at the edge of the waveguide where its intensity is measured. The coupling condition for a first order diffraction linear grating in air is n* = nair sin + /; where n*(nf ,t,) is the effective index of the waveguide mode, nf is the silicon oxynitride film refractive index, t is the film thickness, is the wavelength, and is the grating pitch. For a waveguide with nf = 1.8, thickness t = 330 nm, grating pitch = 1.0 m, and center wavelength ∼ 1.45 m, coupling angles when the grating is exposed to aqueous medium occur near normal incidence at = −2◦ for TE and = −4.2◦ for the TM mode. A typical FWHM of the resonance curve is 1.5–2 nm. The intensity of the light coupled into the waveguide is measured in the TWIST apparatus as function of the laser wavelength with the waveguide device at a fixed angle value. Light emerging from the waveguide is collected by a multimode fiber aligned at the edge of the sample and connected to a photodiode. The measured intensity values describe a bell-shaped curve whose wavelength position shifts according to the refractive index values at the grating surface. If the surface is properly functionalized, this shift is the signature for specific molecular binding. Both TE and TM peaks shift, but typically by different amounts because the two modes have different dependences on refractive index. The waveguide thickness was selected to achieve greater sensitivity with the TM mode, as calculated using the approach described in Refs. [15,16]. 2.2. Device fabrication Grating-integrated optical waveguide devices were fabricated as described previously [13,17]. Briefly, the silicon oxynitride films were deposited in a capacitively coupled PECVD system (Oxford PlasmaLab80® Plus) at rf frequencies of 100 kHz [18]. The refractive index of the film can be varied by gas composition and for this study we used films with a core index of nf = 1.8 and
thickness of t = 330 nm for single mode operation. The waveguide films were deposited either on pyrex glass wafers or on oxidized silicon wafers (with a 10 m oxide layer acting as lower cladding). A grating pattern was integrated with the slab waveguide using thermal nanoimprint lithography with a commercial replica grating as a template followed by dry etching. Wafers were diced (typical dimensions 40 mm × 30 mm) before imprinting and etching. Imprinting was performed on a die bonder tool (Smart Equipment Technology FC-150) at T = 160 ◦ C and 40 kg of force. Reactive ion etching (Oxford PlasmaLab 800 Plus) with CHF3 /Ar/O2 was used to transfer the pattern. In order to improve light coupling efficiency, a grating pattern with a shallower corrugation depth than previously used was selected to reduce the perturbation to the waveguide light transmission caused by the grating. Holographic replica gratings (3-4067, Optometrics) with grooves on a = 1 m pitch, corrugation of 100 nm and 12.6 mm × 12.6 mm area were used as low-cost, commercially available stamps and 100 nm thick thermal imprint polymer (MR-8010, Microresist) was used as sacrificial resist. An etch time of 5 min resulted in grooves of hg = 80 nm depth, as measured by AFM (Veeco). No polishing of the waveguide edges was required. The grating depth hg relative to film thickness t in our devices is larger than values reported in other works (hg is 25% of the thickness while values of 8–12% have been reported [9,19]). The grating depth is known to affect the dependence of sensitivity on waveguide film thickness [20]. The thickness of the waveguide film in this work (t = 330 nm) was selected based on the simple thin grating approximation. Our fabrication approach enables control of the grating depth by the etch time used to transfer the grating pattern. We tested a device with half the grating depth (hg = 40 nm), and observed no significant resonance peak broadening nor change in sensing performance. 2.3. Sensor cartridge Devices were integrated in fluidic cartridges as illustrated in Fig. 2(a). A 120 m thick spacer defining the flow channel was cut out of a SecureSealTM double-sided adhesive sheet (Grace Biolabs). The flow channel dimensions are 24 mm × 4 mm, with the length aligned parallel to the groove direction (and waveguide edge). The width of the channel was selected to allow convenient optical access to a 1 mm wide laser beam. The laser beam is incident near the edge of the grating pattern to avoid light outcoupling back through the grating rather than propagation of the guided mode, and a good portion 12.6 mm × 12.6 mm imprinted area is not needed. The 40 mm × 14 mm glass lid contains two 1 mm holes to which NanoPortTM connectors (N-333, Idex) were attached for tubing connections. The flow is parallel to the grooves of the grating and to the waveguide edge being measured. An acetone-soluble sealant (nail polish) is hand-painted on the chamber edges. This assembly (Fig. 2(b)) used with 1/32 in. ID Tygon® tubing provides leak-free operation at flows up to hundreds of l/min with no degradation observed over several days of tests. The cartridge is conveniently disassembled in acetone and re-used after surface-binding analyte tests: the adhesive spacer is discarded, the NanoPortTM assembly is cleaned in acetone and the device is cleaned with acetone followed by piranha (4 H2 SO4 :1 H2 O2 ). Liquid flow is controlled by a peristaltic pump in aspiration mode as shown in Fig. 2(c). The inlet tubing is flexible 20 cm long 1/32 in. ID Tygon® tubing, connected using 1 in. long pieces of rigid PTFE tubing. The use of an inline injection valve is avoided with the simple arrangement of keeping the sample reservoir physically higher than the device, so that gravitational pressure does not cause introduction of air pockets in the sample line when moving the inlet
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Fig. 2. (a) Illustration showing exploded view of the sensor cartridge exposing its three elements, the device, the double-side adhesive layer and glass lid with glued tubing connectors. (b) Photograph of the assembled cartridge. (c) Schematic of fluidic operation.
tubing between different reservoirs. The inlet tubing length corresponds to a sample volume of ∼100 l, while the volume contained in the device chamber (dimensions ∼24 mm × 4 mm × 0.12 mm) is ∼12 l.
signal/background ratio as compared to the photodetector; but the signal to noise ratio remained high enough for sensing measurements, ranging from 40 to 250.
2.4. Optical reader components
2.4.3. Data acquisition and processing A custom four-input data acquisition (DAQ) board was designed and implemented to digitize the signal of four photodiodes. The DAQ consists of a multi-stage amplifier followed by 16-bit analogto-digital converter as shown in Fig. 3(a). The first stage amplifier uses two current amplifiers which swing in opposite directions around a bias voltage to utilize the full range of the voltage supply. The bias voltage is established at the midpoint of the voltage supply using a simple voltage divider and a unity gain amplifier. A second stage amplifies the voltage difference generated by the current amplifiers and connects to the A/D converter. The outputs of the A/D converter are read by a microcontroller and made available to the user via a USB interface. The board has four independent simultaneous input readings and power for the electronics is provided via USB, but can be easily transferred to batteries as all electronics operate from a single 3.3 V supply. LabVIEW code was developed to integrate the board readings with the software interface controlling the laser wavelength and performing data processing. Photodiode intensity spectra are acquired at the smallest wavelength step of the laser, 0.2 nm and the resonance peak position is determined using a centroid method [21] on linearly interpolated data.
2.4.1. Laser source A distributed feedback (DFB) laser module designed for fiberoptic telecommunications (model TL2020-C, Santur) provides linearly polarized fiber-coupled 20 mW power output with a wavelength range over the 36 nm C-band (1528.77–1563.05 nm) and channel spacing of 25 GHz (approximately 0.2 nm). This DWDM (dense wavelength division multiplexing) laser offers the smallest form factor (76 mm × 51 mm × 13 mm) we could find on the market for such a wide tuning range and it is relatively inexpensive (∼$1000 USD for a single unit). The manufacturer provided an evaluation board to computer control the laser via RS232. The laser tuning range (36 nm) is larger than thermally tuned lasers (such as VCSEL, tuning range ∼2 nm) and adequate to interrogate a coupling peak with FWHM of 1.5–2 nm. This tuning range also provides tolerance for mechanical misalignment of the sensor. This is a convenient feature for a system to be used in the field where the waveguide devices are to be manually positioned on a holder at a fixed angle. For our experimental parameters, an angle variation corresponds to a wavelength shift (nm) ∼ 18 (◦ ). This means that a mechanical angular misalignment of as little as 0.1◦ results in a coupling peak position shifted by 1.8 nm, which is, however, well within the tuning range. 2.4.2. Photodetection The waveguided laser mode is outcoupled at the edge of the waveguide slab and collected by a 400 m diameter optical fiber (Thorlabs); the output fiber alignment has a large position tolerance (∼ mm). The fiber is coupled to a germanium photodiode (Thorlabs, SM05PD6A) with a 3 mm active area and no active circuitry. Using a multimeter (Keithley 2000), we validated the photodiode performance against a photodetector (Agilent 81623) and verified that W peak light intensity corresponds to A peak current. The photodiode dark current produced a slight (20%) degradation of the
2.4.4. Multichannel operation The use of an integrated grating sensor device and fiber-coupled photodiodes as photodetectors enables a straightforward approach to obtain multiple sensors on the same device. Spatial regions of the same grating parallel to the waveguide edge can become separate sensors if simultaneously illuminated and probed by multiple optical fibers. Simultaneous illumination of multiple regions along the 12 mm long grating is achieved by laser beam expansion. The laser fiber output after having been collimated to a 1 mm diameter beam is unidimensionally expanded with a pair of cylindrical lenses with focal length f at f1 + f2 distance (Thorlabs, LJ1567L1C, f1 = 100 mm, and LK1836L2-C, f2 = −9.7 mm). Two 400 m core
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3. Materials and methods 3.1. Volume sensing In order to quantify the sensor sensitivity and resolution to changes of bulk refractive index, solutions of different concentrations of NaCl and isopropanol in deionized water were used. The refractive indices of the solutions were obtained from data tables [22] for isopropanol dilution up to 1%, then by calculation to values diluted to 0.2%. For NaCl solutions, 0.0018 RIU (refractive index units) shift per mass % was used [23]. The refractive index solutions were injected in the fluidic cartridge after a deionized water baseline was equilibrated for 2–4 min and a measurement taken; the cartridge was then thoroughly rinsed with water before the next solution was introduced. 3.2. Surface sensing The device surface was cleaned using piranha solution (HSO4 70% and H2 O2 30%) for 1 h and rinsed with deionized water. The surface was aminofunctionalized by soaking in a 3% solution of APTMS (aminopropytrimethoxy silane, Gelest SIA0611.0) in methanol with 2% water for 1 h, rinsed in methanol and water, dried under a gentle N2 flow and stored in a desiccator until use. A 5% solution of glutaraldehyde (Sigma) in deionized water was applied to activate the surface, with incubation time of 30 min and rinsed with deionized water. Rabbit anti-goat IgG antibody (Thermoscientific 31105, stock concentration 2.4 mg/ml) was serially diluted in PBS buffer (PBS (phosphate buffer saline, 0.1 M sodium phosphate and 0.15 M NaCl). Increasing concentrations of IgG were then injected on a device cartridge at flow rate of 12 l/min. The resonance shift was monitored as a 150 l sample plug flowed across the activated surface. After 5 min of monitoring under flow, the flow was stopped and the resonance shift was recorded until saturation, after which the next solution was introduced. The characterization measurements reported in this paper were performed for convenience on an earlier and bulkier version of the rotary stage than the one shown in Fig. 3 but with the optical and fluidic components (DWDM, laser, photodiode, DAQ and fluidics) as shown. The use of the smaller stage does not change the device performance. Fig. 3. (a) Circuit diagram of the custom DAQ. The photodiode current (IPD ) is amplified by two current amplifiers which swing in opposite directions around a bias voltage creating a voltage difference of 2IPD R1 at the input of a differential amplifier. The differential amplifier provides a gain of R3 /R2 and connects to the A/D converter. (b) Photograph of a benchtop prototype including all the components for the TWIST instrument with two sensor outputs. Scale bar is 2.5 cm. Food coloring is used to highlight the flow path.
fibers are arranged at 6 mm interdistance to collect outcoupled radiation.
2.4.5. Optical reader size The fiber arrangement and system components are shown in Fig. 3(b). The laser module is plugged into a large-footprint PCB board for power and communication purposes and protected in an enclosure, however, the PCB control board can ultimately be reduced in size to the same footprint of the laser. The prototype shown in Fig. 3(b) is not packaged in the most compact possible format, as this is the object of ongoing effort. A SolidWorks model of the optical components (not shown) indicates that, when consolidated, the optical reader fits within the volume of a handheld device 100 mm (l) × 50 mm (w) × 100 mm (h) excluding fluidic pump.
4. Results 4.1. Volumetric sensing Grating devices were exposed to aqueous solutions with refractive indices ranging from 5 × 10−5 RIU to 0.012 RIU in order to determine volume refractive index sensitivity and limit of detection of the system. The sensitivity (S) is defined as wavelength shift per refractive index unit. After preliminary evaluation of both TE and TM mode, we found that the sensitivity of the TM mode was roughly twice as large for the TE mode so we focused the remainder of the analysis reported in this paper on TM modes. The measured wavelength shift is linearly dependent on the refractive index of the solution as expected (Fig. 4), and the shift values for the NaCl and IPA solutions are in good agreement. The range of refractive index covered by this calibration up to 0.012 RIU is most relevant for biomedical detection, corresponding to an overall wavelength shift of 1.5 nm, and is neither limited by the laser wavelength range nor by the sensing principle. The sensitivity of the device is obtained by the slope of a linear fit of the data, indicating S = 141 nm/RIU. The chip to chip average value for sensitivity was 142 ± 6 nm/RIU (n = 4). Reproducibility expressed as coefficient of variation (CV) was CV = 4.4%.
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Fig. 4. (a) Wavelength shift as function of refractive index of the solutions measured. The slope of linear fit provides the sensitivity S value of 142 nm/RIU. Inset shows the wavelength shift measured for a baseline to illustrate the system noise. (b) A magnification of the data in the square box in (a). Diamonds corresponds to isopropanol and open squares to NaCl solutions.
Adopting established conventions for refractive index sensing transducers [2], the limit of detection LOD is defined as R/S, where S is sensitivity and R is sensor resolution. The resolution is expressed in terms of standard deviation of noise of the sensor output. Typical values of standard deviation of baseline in our measurements was on the order of 1–2 pm, depending on the number of measurements; the inset of Fig. 4(a) shows over 20 measurements performed on a stabilized device exposed to an aqueous solution. The standard deviation of the data set is = 1.5 pm which results in a LOD of 1 × 10−5 RIU. This result improves over previously reported volumetric sensing in this configuration because of a combination of higher signal to noise (achieved by improving the grating corrugation depth) and optimized centroid calculation of the wavelength shift.
0.2 to 200 g/ml) the wavelength shift increased by a factor 20 (0.08–0.15 nm). The typical acquisition time of one wavelength spectrum is 2–3 min, and it is determined by the wavelength switching speed of the DWDM laser, which is not designed for wavelength sweeping and requires a few seconds per wavelength. As a trade-off for the selection of a highly compact tunable source, the system has a limited data acquisition speed. The speed is adequate for affinitybased detection where sensor response is typically monitored over 10–30 min. The lowest protein concentration whose binding was resolved was 0.2 g/ml, which corresponds to an approximate surface mass density of 20 pg/mm2 as estimated from the fluidic chamber area.
4.2. Surface sensing
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The most straightforward test for probing the surface sensing capability of the system is the measurement of non-specific protein adsorption on an activated sensor surface. Binding of increasing concentrations of IgG molecules on an amine-reactive glutaraldehyde-coated grating surface was used to characterize surface sensing. The IgG solutions were loaded in the flow system as 150 l sample plugs of increasing concentration. The sample flowed over the sensor surface at a rate of 12 l/min for 5 min while monitoring the resonance shift. Then the flow was stopped and the resonance was monitored in static conditions until a saturation value was reached. The volume of the sensor chamber is approximately 10 l, much smaller than the 150 l sample plug. This configuration ensures minimal dilution of the sample under test. Also, multiple 10 l volumes of the sample can be sequentially measured from the same sample plug. Fig. 5 shows resonance shift as function of time for surface binding of IgG samples of different concentrations. A sensor response in the form of resonance wavelength increase was observed within the first few minutes following injection of the sample in the chamber, and it was tracked for approximately 20 min. In order to validate the repeatability of the observed shift, a second injection from the same sample plug was measured and it also caused an upshift indicating remaining availability of surface binding sites. Sensor responses in the form of wavelength shifts were observed for a range of increasing protein concentrations until a 1 mg/ml concentration where no shift was registered indicating saturation of binding sites. As typically observed in surface binding-based systems [3,10], the sensor response for equilibrium adsorption of protein increased logarithmically with the concentration. The graphs in Fig. 5 indicate that for protein concentration spanning 3 orders of magnitude (from
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The surface sensing measurements were obtained in stopped-flow conditions. The mass transport limitation of stopped-flow measurement can be lifted and signal from diluted samples can be maximized by acquiring data under continuous flow with the appropriate Peclet number to replenish the analyte near the transducer surface as reported by others [8,10,24] as well as adoption of clever valving configurations to introduce sample to the sensing area without dispersion as recently reported [25]. 4.3. Multiplexed operation The simultaneous temporal response from two sensors in response to different refractive index samples is shown in Fig. 6. The two sensors were exposed to the same solution flowing in the cartridge: the baseline used for this test was phosphate buffer saline (PBS) and index changes were induced by solution of NaCl and HCl in deionized water with lower and higher refractive index. The two sensors’ coefficient of variation in response to the same change in refractive index is no larger than 15%. Furthermore, the sensors track each other in temporal drift, indicating that this configuration is effective for on-chip referencing in binding experiments where the two grating regions probed by the two detectors have different functionalization. 5. Discussion We demonstrated that the TWIST sensing system can be implemented with a small footprint and cost-effective components and good performance. The TWIST platform offers a combination of advantages from both optical reader components and device structure over other compact optical sensors [7,10], and these features will ultimately produce a cost-effective and user-friendly handheld instrument with appropriate sensitivity for a variety of assays. The use of the edge-detection approach avoids the need for folded optical paths for the light output collection in reflection or transmission modes and can therefore reduce the system to hand-held size. Edge-detection is enabled by the long propagation length of coupled modes in optical waveguides and it is not available in competing sensing approaches such as surface plasmon resonance. The use of an input grating coupler streamlines the chip microfabrication steps (as compared to approaches requiring multiple gratings and multilayer fabrication) and avoids the use of prisms [11]. The selection of high optical transmission SiON waveguides affords tolerance in the selection of the optical window of operation. Employing a largely tunable DWDM laser in the 1.5 m wavelength regions relaxes the requirement of angular alignment of the chip which enables easy mechanical placement of the chip in the
read-out apparatus. The system is insensitive to ambient light. The fiber-based outcoupled light collection is alignment-tolerant and does not require optical elements such as lenses. The most stringent requirements of optical interrogation of the device (wavelength stability, large tuning range) are satisfied by commercial compact solid-state lasers developed for the telecommunication industry, so continued performance enhancement and cost reduction are anticipated. In particular the development of fast sweeping broadband sources in the infrared for emerging biomedical techniques such as optical coherence tomography could readily enhance our data acquisition speed. The illumination occurring through the transparent substrate provides flexibility in the arrangement of fluidics over the transducer surface, as opposed to approaches that require illumination through the cover window, and therefore de-couples the liquid handling constraints from those of the optical reader. The mass surface sensitivity obtained for protein binding (20 pg/mm2 ) is comparable to figures reported recently by other groups [26] and can be improved by appropriate analyte fluidic delivery and by optimizing the grating-waveguide thickness structure. An equivalent layer approximation calculation of sensitivity for a system with grating thickness hg = 80 nm was performed with the approach described in Ref. [26], and it indicated a different optimal film thickness value (t = 400 nm) for the TM mode. We note, however, that the relatively long wavelength provides penetration depth suitable for detection of large analytes such as bacteria at the expense of sensitivity. It is unlikely that optimized performance will enable the sub pg/mm2 sensitivity reported by more complex approaches not compatible with our effort of integration into a hand-held package [8,27]. In order to meet the stringent detection requirements for clinical applications such as detection of virus and bacteria with a fieldable system, we plan to move beyond the direct binding assay approach, and exploit sandwich assay enhancement as well as explore on-chip electrokinetic sample concentration approaches, such as the one which was recently demonstrated with compatibility for grating-integrated sensors [28]. 6. Conclusion The design, assembly and characterization of a biosensor instrument based on the input grating coupler have been presented. A system based on wavelength interrogation and edge detection of guided modes has been developed with compact optical components including a DWDM laser source and photodiode read-outs. We have determined the volume sensitivity and limit of detection of the system as 142 nm/RIU and 1 × 10−5 , respectively, and we have measured surface mass sensing to protein solution as low as 200 ng/ml. We presented an approach to obtain multiple sensors on the same device and characterized the reproducibility of a twosensor system. The ultimate implementation of this biosensor as a field instrument for applications such as infectious disease diagnosis at the point-of-care will require the packaging of the optical and fluid transport components in a hand-held instrument. Acknowledgement This work was partially supported by DARPA through SPAWARSYSCEN Grant Number N66001-04-1-8933. References [1] P.V. Lambeck, Integrated optical sensors for the chemical domain, Measurement Science and Technology 17 (2006) R93–R116. [2] J. Homola, Surface plasmon resonance sensors for detection of chemical and biological species, Chemical Reviews 108 (2008) 462–493. [3] N. Adanyi, I.A. Levkovets, S. Rodriguez-Gil, A. Ronald, M. Varadi, I. Szendrõ, Development of immunosensor based on OWLS technique for determining
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Biographies Sonia Grego received her PhD in physics in 1999 from the University of Copenhagen, Denmark. Her post-doctoral research in biophysics was focused on the development of optical systems for biological studies (laser tweezers) and optical microscopy. As a scientist at RTI International’s Center for Materials and Electronic Technologies, her research interest include microfabricated devices for optoelectronics and life science applications, with emphasis on biosensing and lab on a chip technologies, photonic nanostructures, and optical MEMS for biomedical imaging. Kristin H. Gilchrist received the bachelor’s degree in biomedical engineering from Vanderbilt University, Nashville, TN, in 1997, and the MS and PhD degrees in electrical engineering from Stanford University, Palo Alto, CA, in 1999 and 2003, respectively. Her doctoral research focused on cell-based sensors utilizing cardiac cells cultured on microelectrode arrays. She is currently a research scientist in the Center for Materials and Electronic Technologies at RTI International. Her research interests include MEMS-based optical and ultrasound medical imaging devices. James B. Carlson received dual BS degrees in mechanical engineering and in electrical engineering from North Carolina State University (1983). His expertise is in embedded system design, ASIC development and transitioning development devices/systems into products. As a scientist at RTI International’s Center for Materials and Electronic Technologies, his efforts include development of mixed mode MEMS control ASICs and low power handheld embedded systems. He is also an adjunct lecturer on Embedded Systems at North Carolina State University. Brian R. Stoner received his BS and MS degrees in electrical engineering and materials science, respectively, from the University of Virginia, and PhD in materials science and engineering from North Carolina State University. His research interests are related to biomedical materials and devices, microelectromechanical systems (MEMS)-based sensors, and photonics. He also is an adjunct professor in the department of electrical and computer engineering at Duke University, where he conducts research on biomaterials, nanostructured materials, and related devices. He holds 23 U.S. patents related to novel microelectronic materials and systems, and has authored or co-authored two book chapters and more than 140 publications.
Contents lists available at SciVerse ScienceDirect
Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb
A compact and multichannel optical biosensor based on a wavelength interrogated input grating coupler Sonia Grego ∗ , Kristin H. Gilchrist, James B. Carlson, Brian R. Stoner Center for Materials and Electronic Technologies, RTI International, 3040 E. Cornwallis Road, Research Triangle Park, NC 27709-2194, USA
a r t i c l e
i n f o
Article history: Received 14 August 2011 Received in revised form 27 October 2011 Accepted 7 November 2011 Available online 25 November 2011 Keywords: Optical waveguide Tunable laser Grating Multichannel sensor Portable biosensor
a b s t r a c t The design, fabrication and characterization of a compact optical biosensor based on an input grating coupler are presented. The sensor device consists of a waveguide-integrated single grating structure. The optical reader is based on the use of a largely tunable, small footprint infrared laser for wavelength interrogation of the sensor and on edge-detection of waveguided modes, resulting in a compact configuration. A multiplexing approach is illustrated and a two-output sensor system is demonstrated. The device is packaged in a microfluidic cartridge for sample handling. The volumetric sensing performance of the system is characterized and a limit of detection of 1 × 10−5 RIU is reported while the surface sensing capability is demonstrated by detection of protein diluted down to 200 ng/ml. © 2011 Elsevier B.V. All rights reserved.
1. Introduction Label-free optical sensing technologies based on detection of molecular binding events have been extensively investigated as a rapid analytical method for biological, medical and environmental applications. These approaches detect biomolecule absorption on the transducer surface as a change in the refractive index at the transducer–liquid interface using evanescent waves in either all dielectric or dielectric-metal devices [1,2]. Multichannel biosensor instruments reading receptor-functionalized chips have been demonstrated as powerful immunosensing tools but are limited to laboratory use due to their large size and complexity. Fielddeployable platforms are highly desirable, given the biosensor capability to quickly detect contaminants in food [3,4] or diagnose infections from a clinical specimen such as blood [5], however, they often provide reduced sensitivity compared to benchtop instruments. In recent years, efforts have been reported to develop portable sensors using fully packaged approaches based on surface plasmons [6–8] or grating couplers [9] or to reduce size by integration of sample handling on-chip (e.g. in ring resonators [10]) yet these systems do not achieve a truly hand-held size or do so at expense of lack of multiplexing or non-scalable device manufacturing [11,12]. We present a grating coupler sensor based on wavelength interrogation and edge detection which allows construction of a
∗ Corresponding author. Tel.: +1 919 248 4181; fax: +1 919 248 1955. E-mail address: [email protected] (S. Grego). 0925-4005/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.snb.2011.11.020
compact and rugged sensor suitable for field use. The approach is referred to as Tunable Wavelength Interrogated Sensor Technology (TWIST) and relies on the use of a largely tunable yet compact infrared wavelength source. We previously fabricated gratingintegrated waveguide devices and characterized their sensing performance at = 1.5 m using benchtop components including a wavelength swept laser source for telecommunication testing [13]. The advantage of a broadband wavelength sweeping source for optical waveguide grating sensors was also demonstrated by others [14] by using a custom-assembled fiber-based laser source implemented with an electronically tunable Fabry–Perot filter and other components. That work demonstrated a large wavelength scan at kHz speed driven by an external voltage source and was realized as a benchtop sensor apparatus. This paper describes a miniaturized system serving as optical reader of the waveguide devices including a new optical source and photodetector as well as multichannel capability. The configuration is designed and the components are selected with the ultimate objective of integration into a hand-held device interrogating disposable sensor chips. Major requisites of the system include the use of compact and inexpensive components and the ability to multiplex sensors on the same device to enable on-chip referencing. The waveguide grating sensor chip was integrated with a fluidic cartridge and the system detection capabilities were characterized in terms of volume refractive index sensing and surface sensing of IgG antibody. The miniaturized system performance is improved compared to our previously reported values to a detection limit of 1 × 10−5 RIU and an on-chip reference configuration is demonstrated.
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Fig. 1. Schematic of the TWIST sensor principle.
2. Tunable wavelength interrogated sensor technology (TWIST) 2.1. Principle of operation TWIST is based on evanescent wave sensing of surface binding analytes using a wavelength-tunable, input grating-coupler configuration. A schematic of the sensing principle is shown in Fig. 1. A collimated laser beam is incident on the grating, coupled into the silicon oxynitride waveguide film, and emerges at the edge of the waveguide where its intensity is measured. The coupling condition for a first order diffraction linear grating in air is n* = nair sin + /; where n*(nf ,t,) is the effective index of the waveguide mode, nf is the silicon oxynitride film refractive index, t is the film thickness, is the wavelength, and is the grating pitch. For a waveguide with nf = 1.8, thickness t = 330 nm, grating pitch = 1.0 m, and center wavelength ∼ 1.45 m, coupling angles when the grating is exposed to aqueous medium occur near normal incidence at = −2◦ for TE and = −4.2◦ for the TM mode. A typical FWHM of the resonance curve is 1.5–2 nm. The intensity of the light coupled into the waveguide is measured in the TWIST apparatus as function of the laser wavelength with the waveguide device at a fixed angle value. Light emerging from the waveguide is collected by a multimode fiber aligned at the edge of the sample and connected to a photodiode. The measured intensity values describe a bell-shaped curve whose wavelength position shifts according to the refractive index values at the grating surface. If the surface is properly functionalized, this shift is the signature for specific molecular binding. Both TE and TM peaks shift, but typically by different amounts because the two modes have different dependences on refractive index. The waveguide thickness was selected to achieve greater sensitivity with the TM mode, as calculated using the approach described in Refs. [15,16]. 2.2. Device fabrication Grating-integrated optical waveguide devices were fabricated as described previously [13,17]. Briefly, the silicon oxynitride films were deposited in a capacitively coupled PECVD system (Oxford PlasmaLab80® Plus) at rf frequencies of 100 kHz [18]. The refractive index of the film can be varied by gas composition and for this study we used films with a core index of nf = 1.8 and
thickness of t = 330 nm for single mode operation. The waveguide films were deposited either on pyrex glass wafers or on oxidized silicon wafers (with a 10 m oxide layer acting as lower cladding). A grating pattern was integrated with the slab waveguide using thermal nanoimprint lithography with a commercial replica grating as a template followed by dry etching. Wafers were diced (typical dimensions 40 mm × 30 mm) before imprinting and etching. Imprinting was performed on a die bonder tool (Smart Equipment Technology FC-150) at T = 160 ◦ C and 40 kg of force. Reactive ion etching (Oxford PlasmaLab 800 Plus) with CHF3 /Ar/O2 was used to transfer the pattern. In order to improve light coupling efficiency, a grating pattern with a shallower corrugation depth than previously used was selected to reduce the perturbation to the waveguide light transmission caused by the grating. Holographic replica gratings (3-4067, Optometrics) with grooves on a = 1 m pitch, corrugation of 100 nm and 12.6 mm × 12.6 mm area were used as low-cost, commercially available stamps and 100 nm thick thermal imprint polymer (MR-8010, Microresist) was used as sacrificial resist. An etch time of 5 min resulted in grooves of hg = 80 nm depth, as measured by AFM (Veeco). No polishing of the waveguide edges was required. The grating depth hg relative to film thickness t in our devices is larger than values reported in other works (hg is 25% of the thickness while values of 8–12% have been reported [9,19]). The grating depth is known to affect the dependence of sensitivity on waveguide film thickness [20]. The thickness of the waveguide film in this work (t = 330 nm) was selected based on the simple thin grating approximation. Our fabrication approach enables control of the grating depth by the etch time used to transfer the grating pattern. We tested a device with half the grating depth (hg = 40 nm), and observed no significant resonance peak broadening nor change in sensing performance. 2.3. Sensor cartridge Devices were integrated in fluidic cartridges as illustrated in Fig. 2(a). A 120 m thick spacer defining the flow channel was cut out of a SecureSealTM double-sided adhesive sheet (Grace Biolabs). The flow channel dimensions are 24 mm × 4 mm, with the length aligned parallel to the groove direction (and waveguide edge). The width of the channel was selected to allow convenient optical access to a 1 mm wide laser beam. The laser beam is incident near the edge of the grating pattern to avoid light outcoupling back through the grating rather than propagation of the guided mode, and a good portion 12.6 mm × 12.6 mm imprinted area is not needed. The 40 mm × 14 mm glass lid contains two 1 mm holes to which NanoPortTM connectors (N-333, Idex) were attached for tubing connections. The flow is parallel to the grooves of the grating and to the waveguide edge being measured. An acetone-soluble sealant (nail polish) is hand-painted on the chamber edges. This assembly (Fig. 2(b)) used with 1/32 in. ID Tygon® tubing provides leak-free operation at flows up to hundreds of l/min with no degradation observed over several days of tests. The cartridge is conveniently disassembled in acetone and re-used after surface-binding analyte tests: the adhesive spacer is discarded, the NanoPortTM assembly is cleaned in acetone and the device is cleaned with acetone followed by piranha (4 H2 SO4 :1 H2 O2 ). Liquid flow is controlled by a peristaltic pump in aspiration mode as shown in Fig. 2(c). The inlet tubing is flexible 20 cm long 1/32 in. ID Tygon® tubing, connected using 1 in. long pieces of rigid PTFE tubing. The use of an inline injection valve is avoided with the simple arrangement of keeping the sample reservoir physically higher than the device, so that gravitational pressure does not cause introduction of air pockets in the sample line when moving the inlet
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Fig. 2. (a) Illustration showing exploded view of the sensor cartridge exposing its three elements, the device, the double-side adhesive layer and glass lid with glued tubing connectors. (b) Photograph of the assembled cartridge. (c) Schematic of fluidic operation.
tubing between different reservoirs. The inlet tubing length corresponds to a sample volume of ∼100 l, while the volume contained in the device chamber (dimensions ∼24 mm × 4 mm × 0.12 mm) is ∼12 l.
signal/background ratio as compared to the photodetector; but the signal to noise ratio remained high enough for sensing measurements, ranging from 40 to 250.
2.4. Optical reader components
2.4.3. Data acquisition and processing A custom four-input data acquisition (DAQ) board was designed and implemented to digitize the signal of four photodiodes. The DAQ consists of a multi-stage amplifier followed by 16-bit analogto-digital converter as shown in Fig. 3(a). The first stage amplifier uses two current amplifiers which swing in opposite directions around a bias voltage to utilize the full range of the voltage supply. The bias voltage is established at the midpoint of the voltage supply using a simple voltage divider and a unity gain amplifier. A second stage amplifies the voltage difference generated by the current amplifiers and connects to the A/D converter. The outputs of the A/D converter are read by a microcontroller and made available to the user via a USB interface. The board has four independent simultaneous input readings and power for the electronics is provided via USB, but can be easily transferred to batteries as all electronics operate from a single 3.3 V supply. LabVIEW code was developed to integrate the board readings with the software interface controlling the laser wavelength and performing data processing. Photodiode intensity spectra are acquired at the smallest wavelength step of the laser, 0.2 nm and the resonance peak position is determined using a centroid method [21] on linearly interpolated data.
2.4.1. Laser source A distributed feedback (DFB) laser module designed for fiberoptic telecommunications (model TL2020-C, Santur) provides linearly polarized fiber-coupled 20 mW power output with a wavelength range over the 36 nm C-band (1528.77–1563.05 nm) and channel spacing of 25 GHz (approximately 0.2 nm). This DWDM (dense wavelength division multiplexing) laser offers the smallest form factor (76 mm × 51 mm × 13 mm) we could find on the market for such a wide tuning range and it is relatively inexpensive (∼$1000 USD for a single unit). The manufacturer provided an evaluation board to computer control the laser via RS232. The laser tuning range (36 nm) is larger than thermally tuned lasers (such as VCSEL, tuning range ∼2 nm) and adequate to interrogate a coupling peak with FWHM of 1.5–2 nm. This tuning range also provides tolerance for mechanical misalignment of the sensor. This is a convenient feature for a system to be used in the field where the waveguide devices are to be manually positioned on a holder at a fixed angle. For our experimental parameters, an angle variation corresponds to a wavelength shift (nm) ∼ 18 (◦ ). This means that a mechanical angular misalignment of as little as 0.1◦ results in a coupling peak position shifted by 1.8 nm, which is, however, well within the tuning range. 2.4.2. Photodetection The waveguided laser mode is outcoupled at the edge of the waveguide slab and collected by a 400 m diameter optical fiber (Thorlabs); the output fiber alignment has a large position tolerance (∼ mm). The fiber is coupled to a germanium photodiode (Thorlabs, SM05PD6A) with a 3 mm active area and no active circuitry. Using a multimeter (Keithley 2000), we validated the photodiode performance against a photodetector (Agilent 81623) and verified that W peak light intensity corresponds to A peak current. The photodiode dark current produced a slight (20%) degradation of the
2.4.4. Multichannel operation The use of an integrated grating sensor device and fiber-coupled photodiodes as photodetectors enables a straightforward approach to obtain multiple sensors on the same device. Spatial regions of the same grating parallel to the waveguide edge can become separate sensors if simultaneously illuminated and probed by multiple optical fibers. Simultaneous illumination of multiple regions along the 12 mm long grating is achieved by laser beam expansion. The laser fiber output after having been collimated to a 1 mm diameter beam is unidimensionally expanded with a pair of cylindrical lenses with focal length f at f1 + f2 distance (Thorlabs, LJ1567L1C, f1 = 100 mm, and LK1836L2-C, f2 = −9.7 mm). Two 400 m core
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3. Materials and methods 3.1. Volume sensing In order to quantify the sensor sensitivity and resolution to changes of bulk refractive index, solutions of different concentrations of NaCl and isopropanol in deionized water were used. The refractive indices of the solutions were obtained from data tables [22] for isopropanol dilution up to 1%, then by calculation to values diluted to 0.2%. For NaCl solutions, 0.0018 RIU (refractive index units) shift per mass % was used [23]. The refractive index solutions were injected in the fluidic cartridge after a deionized water baseline was equilibrated for 2–4 min and a measurement taken; the cartridge was then thoroughly rinsed with water before the next solution was introduced. 3.2. Surface sensing The device surface was cleaned using piranha solution (HSO4 70% and H2 O2 30%) for 1 h and rinsed with deionized water. The surface was aminofunctionalized by soaking in a 3% solution of APTMS (aminopropytrimethoxy silane, Gelest SIA0611.0) in methanol with 2% water for 1 h, rinsed in methanol and water, dried under a gentle N2 flow and stored in a desiccator until use. A 5% solution of glutaraldehyde (Sigma) in deionized water was applied to activate the surface, with incubation time of 30 min and rinsed with deionized water. Rabbit anti-goat IgG antibody (Thermoscientific 31105, stock concentration 2.4 mg/ml) was serially diluted in PBS buffer (PBS (phosphate buffer saline, 0.1 M sodium phosphate and 0.15 M NaCl). Increasing concentrations of IgG were then injected on a device cartridge at flow rate of 12 l/min. The resonance shift was monitored as a 150 l sample plug flowed across the activated surface. After 5 min of monitoring under flow, the flow was stopped and the resonance shift was recorded until saturation, after which the next solution was introduced. The characterization measurements reported in this paper were performed for convenience on an earlier and bulkier version of the rotary stage than the one shown in Fig. 3 but with the optical and fluidic components (DWDM, laser, photodiode, DAQ and fluidics) as shown. The use of the smaller stage does not change the device performance. Fig. 3. (a) Circuit diagram of the custom DAQ. The photodiode current (IPD ) is amplified by two current amplifiers which swing in opposite directions around a bias voltage creating a voltage difference of 2IPD R1 at the input of a differential amplifier. The differential amplifier provides a gain of R3 /R2 and connects to the A/D converter. (b) Photograph of a benchtop prototype including all the components for the TWIST instrument with two sensor outputs. Scale bar is 2.5 cm. Food coloring is used to highlight the flow path.
fibers are arranged at 6 mm interdistance to collect outcoupled radiation.
2.4.5. Optical reader size The fiber arrangement and system components are shown in Fig. 3(b). The laser module is plugged into a large-footprint PCB board for power and communication purposes and protected in an enclosure, however, the PCB control board can ultimately be reduced in size to the same footprint of the laser. The prototype shown in Fig. 3(b) is not packaged in the most compact possible format, as this is the object of ongoing effort. A SolidWorks model of the optical components (not shown) indicates that, when consolidated, the optical reader fits within the volume of a handheld device 100 mm (l) × 50 mm (w) × 100 mm (h) excluding fluidic pump.
4. Results 4.1. Volumetric sensing Grating devices were exposed to aqueous solutions with refractive indices ranging from 5 × 10−5 RIU to 0.012 RIU in order to determine volume refractive index sensitivity and limit of detection of the system. The sensitivity (S) is defined as wavelength shift per refractive index unit. After preliminary evaluation of both TE and TM mode, we found that the sensitivity of the TM mode was roughly twice as large for the TE mode so we focused the remainder of the analysis reported in this paper on TM modes. The measured wavelength shift is linearly dependent on the refractive index of the solution as expected (Fig. 4), and the shift values for the NaCl and IPA solutions are in good agreement. The range of refractive index covered by this calibration up to 0.012 RIU is most relevant for biomedical detection, corresponding to an overall wavelength shift of 1.5 nm, and is neither limited by the laser wavelength range nor by the sensing principle. The sensitivity of the device is obtained by the slope of a linear fit of the data, indicating S = 141 nm/RIU. The chip to chip average value for sensitivity was 142 ± 6 nm/RIU (n = 4). Reproducibility expressed as coefficient of variation (CV) was CV = 4.4%.
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Adopting established conventions for refractive index sensing transducers [2], the limit of detection LOD is defined as R/S, where S is sensitivity and R is sensor resolution. The resolution is expressed in terms of standard deviation of noise of the sensor output. Typical values of standard deviation of baseline in our measurements was on the order of 1–2 pm, depending on the number of measurements; the inset of Fig. 4(a) shows over 20 measurements performed on a stabilized device exposed to an aqueous solution. The standard deviation of the data set is = 1.5 pm which results in a LOD of 1 × 10−5 RIU. This result improves over previously reported volumetric sensing in this configuration because of a combination of higher signal to noise (achieved by improving the grating corrugation depth) and optimized centroid calculation of the wavelength shift.
0.2 to 200 g/ml) the wavelength shift increased by a factor 20 (0.08–0.15 nm). The typical acquisition time of one wavelength spectrum is 2–3 min, and it is determined by the wavelength switching speed of the DWDM laser, which is not designed for wavelength sweeping and requires a few seconds per wavelength. As a trade-off for the selection of a highly compact tunable source, the system has a limited data acquisition speed. The speed is adequate for affinitybased detection where sensor response is typically monitored over 10–30 min. The lowest protein concentration whose binding was resolved was 0.2 g/ml, which corresponds to an approximate surface mass density of 20 pg/mm2 as estimated from the fluidic chamber area.
4.2. Surface sensing
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The most straightforward test for probing the surface sensing capability of the system is the measurement of non-specific protein adsorption on an activated sensor surface. Binding of increasing concentrations of IgG molecules on an amine-reactive glutaraldehyde-coated grating surface was used to characterize surface sensing. The IgG solutions were loaded in the flow system as 150 l sample plugs of increasing concentration. The sample flowed over the sensor surface at a rate of 12 l/min for 5 min while monitoring the resonance shift. Then the flow was stopped and the resonance was monitored in static conditions until a saturation value was reached. The volume of the sensor chamber is approximately 10 l, much smaller than the 150 l sample plug. This configuration ensures minimal dilution of the sample under test. Also, multiple 10 l volumes of the sample can be sequentially measured from the same sample plug. Fig. 5 shows resonance shift as function of time for surface binding of IgG samples of different concentrations. A sensor response in the form of resonance wavelength increase was observed within the first few minutes following injection of the sample in the chamber, and it was tracked for approximately 20 min. In order to validate the repeatability of the observed shift, a second injection from the same sample plug was measured and it also caused an upshift indicating remaining availability of surface binding sites. Sensor responses in the form of wavelength shifts were observed for a range of increasing protein concentrations until a 1 mg/ml concentration where no shift was registered indicating saturation of binding sites. As typically observed in surface binding-based systems [3,10], the sensor response for equilibrium adsorption of protein increased logarithmically with the concentration. The graphs in Fig. 5 indicate that for protein concentration spanning 3 orders of magnitude (from
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Time (min) Fig. 6. Response of two sensor regions measured simultaneously with two detectors while grating is exposed to medium changes. Baseline is PBS buffer. HCl and NaCl solutions in deionized water are as indicated.
The surface sensing measurements were obtained in stopped-flow conditions. The mass transport limitation of stopped-flow measurement can be lifted and signal from diluted samples can be maximized by acquiring data under continuous flow with the appropriate Peclet number to replenish the analyte near the transducer surface as reported by others [8,10,24] as well as adoption of clever valving configurations to introduce sample to the sensing area without dispersion as recently reported [25]. 4.3. Multiplexed operation The simultaneous temporal response from two sensors in response to different refractive index samples is shown in Fig. 6. The two sensors were exposed to the same solution flowing in the cartridge: the baseline used for this test was phosphate buffer saline (PBS) and index changes were induced by solution of NaCl and HCl in deionized water with lower and higher refractive index. The two sensors’ coefficient of variation in response to the same change in refractive index is no larger than 15%. Furthermore, the sensors track each other in temporal drift, indicating that this configuration is effective for on-chip referencing in binding experiments where the two grating regions probed by the two detectors have different functionalization. 5. Discussion We demonstrated that the TWIST sensing system can be implemented with a small footprint and cost-effective components and good performance. The TWIST platform offers a combination of advantages from both optical reader components and device structure over other compact optical sensors [7,10], and these features will ultimately produce a cost-effective and user-friendly handheld instrument with appropriate sensitivity for a variety of assays. The use of the edge-detection approach avoids the need for folded optical paths for the light output collection in reflection or transmission modes and can therefore reduce the system to hand-held size. Edge-detection is enabled by the long propagation length of coupled modes in optical waveguides and it is not available in competing sensing approaches such as surface plasmon resonance. The use of an input grating coupler streamlines the chip microfabrication steps (as compared to approaches requiring multiple gratings and multilayer fabrication) and avoids the use of prisms [11]. The selection of high optical transmission SiON waveguides affords tolerance in the selection of the optical window of operation. Employing a largely tunable DWDM laser in the 1.5 m wavelength regions relaxes the requirement of angular alignment of the chip which enables easy mechanical placement of the chip in the
read-out apparatus. The system is insensitive to ambient light. The fiber-based outcoupled light collection is alignment-tolerant and does not require optical elements such as lenses. The most stringent requirements of optical interrogation of the device (wavelength stability, large tuning range) are satisfied by commercial compact solid-state lasers developed for the telecommunication industry, so continued performance enhancement and cost reduction are anticipated. In particular the development of fast sweeping broadband sources in the infrared for emerging biomedical techniques such as optical coherence tomography could readily enhance our data acquisition speed. The illumination occurring through the transparent substrate provides flexibility in the arrangement of fluidics over the transducer surface, as opposed to approaches that require illumination through the cover window, and therefore de-couples the liquid handling constraints from those of the optical reader. The mass surface sensitivity obtained for protein binding (20 pg/mm2 ) is comparable to figures reported recently by other groups [26] and can be improved by appropriate analyte fluidic delivery and by optimizing the grating-waveguide thickness structure. An equivalent layer approximation calculation of sensitivity for a system with grating thickness hg = 80 nm was performed with the approach described in Ref. [26], and it indicated a different optimal film thickness value (t = 400 nm) for the TM mode. We note, however, that the relatively long wavelength provides penetration depth suitable for detection of large analytes such as bacteria at the expense of sensitivity. It is unlikely that optimized performance will enable the sub pg/mm2 sensitivity reported by more complex approaches not compatible with our effort of integration into a hand-held package [8,27]. In order to meet the stringent detection requirements for clinical applications such as detection of virus and bacteria with a fieldable system, we plan to move beyond the direct binding assay approach, and exploit sandwich assay enhancement as well as explore on-chip electrokinetic sample concentration approaches, such as the one which was recently demonstrated with compatibility for grating-integrated sensors [28]. 6. Conclusion The design, assembly and characterization of a biosensor instrument based on the input grating coupler have been presented. A system based on wavelength interrogation and edge detection of guided modes has been developed with compact optical components including a DWDM laser source and photodiode read-outs. We have determined the volume sensitivity and limit of detection of the system as 142 nm/RIU and 1 × 10−5 , respectively, and we have measured surface mass sensing to protein solution as low as 200 ng/ml. We presented an approach to obtain multiple sensors on the same device and characterized the reproducibility of a twosensor system. The ultimate implementation of this biosensor as a field instrument for applications such as infectious disease diagnosis at the point-of-care will require the packaging of the optical and fluid transport components in a hand-held instrument. Acknowledgement This work was partially supported by DARPA through SPAWARSYSCEN Grant Number N66001-04-1-8933. References [1] P.V. Lambeck, Integrated optical sensors for the chemical domain, Measurement Science and Technology 17 (2006) R93–R116. [2] J. Homola, Surface plasmon resonance sensors for detection of chemical and biological species, Chemical Reviews 108 (2008) 462–493. [3] N. Adanyi, I.A. Levkovets, S. Rodriguez-Gil, A. Ronald, M. Varadi, I. Szendrõ, Development of immunosensor based on OWLS technique for determining
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Biographies Sonia Grego received her PhD in physics in 1999 from the University of Copenhagen, Denmark. Her post-doctoral research in biophysics was focused on the development of optical systems for biological studies (laser tweezers) and optical microscopy. As a scientist at RTI International’s Center for Materials and Electronic Technologies, her research interest include microfabricated devices for optoelectronics and life science applications, with emphasis on biosensing and lab on a chip technologies, photonic nanostructures, and optical MEMS for biomedical imaging. Kristin H. Gilchrist received the bachelor’s degree in biomedical engineering from Vanderbilt University, Nashville, TN, in 1997, and the MS and PhD degrees in electrical engineering from Stanford University, Palo Alto, CA, in 1999 and 2003, respectively. Her doctoral research focused on cell-based sensors utilizing cardiac cells cultured on microelectrode arrays. She is currently a research scientist in the Center for Materials and Electronic Technologies at RTI International. Her research interests include MEMS-based optical and ultrasound medical imaging devices. James B. Carlson received dual BS degrees in mechanical engineering and in electrical engineering from North Carolina State University (1983). His expertise is in embedded system design, ASIC development and transitioning development devices/systems into products. As a scientist at RTI International’s Center for Materials and Electronic Technologies, his efforts include development of mixed mode MEMS control ASICs and low power handheld embedded systems. He is also an adjunct lecturer on Embedded Systems at North Carolina State University. Brian R. Stoner received his BS and MS degrees in electrical engineering and materials science, respectively, from the University of Virginia, and PhD in materials science and engineering from North Carolina State University. His research interests are related to biomedical materials and devices, microelectromechanical systems (MEMS)-based sensors, and photonics. He also is an adjunct professor in the department of electrical and computer engineering at Duke University, where he conducts research on biomaterials, nanostructured materials, and related devices. He holds 23 U.S. patents related to novel microelectronic materials and systems, and has authored or co-authored two book chapters and more than 140 publications.