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Optofluidics in Bio-Chemical Analysis

Yunbo Guo and Xudong Fan

Department of Biomedical Engineering, University of Michigan, 1101 Beal Ave, Ann Arbor, MI 48109-2099, USA

ABSTRACT Optofluidics organically integrates microfluidics and photonics and is an emerging technology in biological and chemical analysis. In this paper, we overview the recent studies in bio-chemical sensing applications of optofluidics. Particularly, we report the research progress in our lab in developing diverse optofluidic devices using two unique configurations: thin-walled capillary based optofluidic ring resonator (OFRR) and multi-hole capillary based optofluidic platforms. The first one has been developed to be OFRR-based label-free biosensor, microfluidic laser based intra-cavity sensors, and on-column optical detectors for micro-gas chromatography (µGC), while the second one has been developed to be optofluidic Fabry-Pérot based label-free biosensor and optofluidic Surface-Enhanced Raman Spectroscopy (SERS) biosensor. All of these devices take advantage of superior fluidic handling capability and high sensitivity, and have been used in detecting various biological and chemical analytes in either liquid or vapor phase.

Keywords: Optofluidics, ring resonators, optical sensor, label-free detection, surface-enhanced Raman scattering, microfluidic laser, microstructured optical fiber, multi-hole capillary, intra-cavity sensor

1. INTRODUCTION

Optofluidics has emerged as a dynamic and rapidly developing research field over the past few years [1], in which optics and microfluidics are synergistically integrated to enhance the performance and function of the microsystems. Optofluidics has found a broad range of applications, as described in several recent reviews in a focus issue of Nature Photonics [2-4]. Particularly, optofluidics is well-suited for biological/chemical detection and analysis in small detection volumes, because it integrates sample preparation and delivery with analysis mechanism. Such integration provides a number of unique characteristics that can be leveraged for biological/chemical analysis. Many optical properties, such as refractive index (RI), fluorescence, Raman scattering, absorption and polarization, can be exploited individually or in combination to generate the sensing signal. Moreover, traditional analytical chemistry technologies such as chromatography and electrophoresis can be adopted to optofluidic devices further increases their functionality in biological/chemical analysis [5, 6]. Furthermore, optofluidic microsystems can also employ optical forces in tandem with microfluidics to trap and manipulate targets [7], thus further enhancing the system’s analytical capabilities. Our lab has been working on diverse optofluidic devices for biological/chemical analysis. These optofluidic devices are mainly based on two configurations: flow-through single-hole capillary and multi-hole capillary with different analytical mechanisms. The first one adopts a thin-wall cylindrical capillary for optofluidic ring resonator (OFRR)[8], which retains inherent microfluidic structures while also possessing excellent optical capabilities of ring resonators. As illustrated in Fig. 1 (a), the OFRR employs a piece of fused silica capillary with a diameter of a few tens to a few hundred micrometers. The capillary serves as the microfluidic channel to deliver liquid or vapor samples, and each capillary circular cross section forms a high finesse ring resonator. The capillary wall is sufficiently thin (< 5 μm) so that the whispering gallery mode (WGM) evanescent field extends on the capillary inner surface (penetration depth ~ 100 nm), interacting with the liquid or vapor sample passing through the capillary or with biomolecules bound to the capillary inner surface. Therefore, the OFRR unique architecture combines the merits of the optical ring resonators and capillary microfluidics, showing high-Q factors, excellent sensitivity, small sample consumption volume, and the excellent fluid handling capability. Originally developed in our lab in early 2006[9], the OFRR has been developed into several optofluidic devices: 1) OFRR-based label-free biosensors for detecting various chemical and biological samples ranging from small molecules such as biotin [10] to large species such as viral particles [11]; 2) OFRR-based surface-enhanced Raman scattering (SERS) sensor for highly sensitive and specific analyte detection [12]; 3) OFRR-based on-column optical detectors for chemical vapor (such as 2,4-Dinitrotoluene (DNT)) detection [13, 14] and micro-gas

Invited Paper

Frontiers in Biological Detection: From Nanosensors to Systems IV, edited by Benjamin L. Miller, Philippe M. Fauchet, Proc. of SPIE Vol. 8212, 82120C

© 2012 SPIE · CCC code: 1605-7422/12/$18 · doi: 10.1117/12.912246

Proc. of SPIE Vol. 8212 82120C-1

Downloaded from SPIE Digital Library on 08 Feb 2012 to 141.213.172.129. Terms of Use: http://spiedl.org/terms

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chromatography (µGC) [15]; 4) OFRR-based microfluidic lasers for low-threshold, tunable lasing [16] [17] and intra-cavity biosensing [18].

Fig.1. Schematic illustrations of (a) a thin-walled capillary based OFRR, (b) multi-hole capillary.

The second optofluidic configuration is multi-hole capillary. As illustrated in Fig. 1 (b), it consists of a bundle of multiple flow-through micro/nano-sized holes. Different from the single-hole capillary, this configuration provides a lot of unique features: 1) its 3-dimensional (3-D) geometry provides large surface-to-volume ratio thus large sensing surface within the detection volume for high sensitivity; 2) its inherent flow-through micro/nano-sized holes (or channels) not only enable convenient sample (in solid, liquid, or vapor phase) delivery, but also increase the analyte capture efficiency due to short diffusion length scales; 3) light can be guided along the capillary for large signal accumulation. Therefore, combing with the start-of-the-art photonic technologies, the multi-hole capillary can be constructed as a variety of devices or systems, opening up a wide range of applications. Since late 2010, our lab have been working on this novel multi-hole capillary, and explored its applications as optofluidic Fabry-Pérot (FP) cavity based flow-through, label-free biosensor [19], optofluidic SERS based biosensor [20], and optofluidic fluorescence immunoassay.

In this paper, we will focus on our recent progress of developing diverse optofluidic devices for sensitive bio-chemical detection and analysis. The optofluidic sensing devices are grouped into two sections based on the capillary configurations shown above. In the first section, the OFRR principle and fabrication method will be presented, followed by the examples of applying OFRR as label-free biosensor for sensitive biomolecular detection, as on-column optical detector for rapid and sensitive vapor detection and µGC analyzer, and as microfluidic laser for intra-cavity biosensing. In the second section, the properties and fabrication of the multi-hole capillary will be investigated, and its applications as optofluidic FP cavity biosensor and SERS biosensor will be also presented.

2. OPTOFLUIDIC RING RESONATORS

2.1 Principle of OFRR In an optofluidic ring resonator (OFRR), the light propagates in the form of whispering gallery modes (WGMs), which result from the total internal reflection of the light along the curved capillary surface. The WGM is a surface mode, and it circulates along the resonator surface and interacts repeatedly with the analytes on its surface through the WGM evanescent field. The WGM spectral position, i.e., resonant wavelength, λ, is related to the refractive index through the resonant condition:

2 effr nm

πλ

⋅ ⋅=

(1)

where r is the ring outer radius, neff is the effective refractive index (RI) experienced by the WGM, and m is an integer that describes the WGM angular momentum. Eq. (1) shows that resonance occurs for wavelengths when an integer multiple of that wavelength matches the circumference. neff changes when the RI near the ring resonator surface is modified due to the RI change of the bulk solution or the capture of target molecules on the surface, which in turn leads to a shift in the WGM spectral position. By directly or indirectly monitoring the WGM spectral shift, it is possible to

(a) (b)



obtain both quantitative and kinetic information about the binding of molecules near the surface. Therefore, the OFRR can be used as label-free optical biosensors for biomolecular detections or vapor detections.

The OFRR sensor sensitivity depends on the interaction between the analytes and the WGM mode via the evanescent field in the liquid core or on the inner surface, which is further highly dependent on the capillary wall thickness. As shown in Fig. 2, the fraction of the mode in the liquid core increases while the wall thickness decreases, which provides various kinds of resonant modes for different sensing applications [10]. For bulk RI detection, large fraction of the mode in the liquid core is preferred, whereas for surface detection, large electric field at the OFRR inner surface is preferred. In addition, for the optofluidic lasing application, the gain medium is filled in the liquid core, thus the wall of the OFRR should be sufficiently thin so that the WGMs of high Q factors (>107) can be exposed in the core to support the low-threshold laser oscillation [16].

Fig. 2. Field distribution of the WGM mode obtained by using the Mie theory to solve a three-layered radial structure. Reprint from Ref. [10] with permission.

2.2 Fabrication of OFRR

Since the thin-walled glass capillary is not commercially available, we have assembled a computer-controlled pulling station in our lab (Fig. 3 (a)), which is capable of fabricating an OFRR up to 50 centimeters in length by rapidly stretching a fused silica preform (outer diameter (OD) normally smaller than 1 mm) under CO2 laser irradiation [8]. As the OFRR wall thickness is critical to its optical properties, attention must be given to the temperature and pulling speed. Therefore, the optimized CO2 power, feed-in speed, and pulling speed are obtained by error and trial for different preforms. With different geometries of capillary preforms, we can fabricate OFRR in a variety of sizes (outer diameter (OD) in the range of 50-150 µm). Fig. 3 (b) shows a microscope image of an OFRR capillary being drawn from a glass capillary preform, and Fig. 3 (c) shows a typical 100 µm OD OFRR with a wall thickness of 5 µm. However, the wall thickness of the pulled OFRR may not be as thin as desired for sensing and lasing purposes. Further reduction of the wall thickness can be carried out with hydrofluoric acid (HF) etching either before pulling or after pulling.

Fig. 3. (a) OFRR fabrication setup, (b)Image of an OFRR made from a capillary preform, (c) SEM image of an OFRR

(b)

(a)

(c)



Laser

Fiber taper

Evanescent field

< 4 µm

Sample flow

WGMs

(A)

(B) (C)

Time

Target injection

Specific binding

OFRR wall

Binding

λ1 λ2

A. U

.

λ1

λ2

2.3 OFRR for label-free biosensors

As discussed in section 2.1, the capillary-based OFRR can be used as label-free biosensors for real-time biomolecular binding measurement. Fig. 4(a)-(c) shows the schematic of the biosensor configuration. Due to extremely high Q (see example shown in Fig. 4(d)), the OFRR biosensor has a very narrow resonance width (< 1 pm) for high detection resolution. Surpassing other ring resonators, the OFRR biosensor with inherent microfluidic channel for convenient analyte delivery, has been developed to be a powerful platform for the detection of a wide range of biomolecules including proteins, DNAs, viruses, bacteria, cells, and clinical related cancer biomarkers [11, 21, 22].

Fig. 4. (a) The capillary circular cross-section forms an optical ring resonator and WGMs are excited by a fiber taper perpendicular to the capillary. The capillary channel serves as the microfluidics for liquid sample delivery. The evanescent field of the WGM at the interior surface interacts with the sample inside the capillary. (b) WGM shifts from λ1 to λ2 in response to the binding of the target molecules to the OFRR interior surface. (c) Sensorgram is obtained by recording the WGM spectral position as a function of time and it provides the kinetic and quantitative information about the biomolecular interaction on the sensor surface. (d) OFRR Q-factor measurement. The transmission spectrum showing a WGM has a Q-factor of 7.5×107.

Furthermore, we investigated the ultimate sensing capability of the capillary-based optofluidic ring resonator with bulk refractive index detection (Fig. 5(a)) and label-free small molecule detection (Fig. 5(b)) [10]. We first optimized our experimental setup, and placed OFRR in an environmentally isolated, thermally controlled copper chamber, which reduced the system noise to about 0.02 pm. In addition, we fabricated OFRRs with the wall thickness to about 1 µm for bulk RI detection and 2 µm for surface mass detection, which maximized the corresponding sensitivity as discussed in section 2.1.

Fig.5 (a) The sensorgram for detection of 2.8×10−7 RIU bulk RI change. (b) The sensorgram for detection of 10 nM biotin binding to streptavidin immobilized on the OFRR surface. Reprint from Ref. [10] with permission

(a)

(b) (c)

(d)

(a) (b)



For bulk RI detection, the OFRR resonant mode had a line width of 8 pm, corresponding to a Q-factor of 1.2×105, which was limited by the intrinsic water absorption at the probe wavelength 980 nm and the OFRR surface roughness caused by the chemical etching. The BRIS was measured to be 570 nm/RIU by using various concentrations of ethanol in deionized water, very close to the simulation result in Fig. 2 (c). The lowest RI change of 2.8×10-7 RIU, is showed in Fig. 5 (a). A 0.15 pm wavelength shift was clearly distinguished when the ethanol-water mixture replaced DI water initially filled in the OFRR. Subsequent DI water rinse brought the mode spectral position back to the baseline, which confirmed that the spectral shift was indeed caused by the RI change in the OFRR core. This result demonstrated the bulk RI sensing capability of the ring resonator in the 10-7 region. Furthermore, the standard deviation, δ, was obtained to be 0.0217 pm, leading to a bulk RI noise equivalent detection limit (NEDL) of 3.8×10-8 RIU, which broke the barrier of 10-7 RIU that label-free biosensors typically encounter.

In surface mass measurement, we chose an OFRR with a wall thickness of about 2 µm. The corresponding BRIS of the mode was approximately 100 nm/RIU and the electric field peaks near the OFRR inner surface, close to the simulation result in Fig. 2 (b). The Q-factor of this mode was about 6×105. Furthermore, we used D-biotin (molecular weight=244 Da) as a model system to test the small molecule detection capability of the OFRR. To detect biotin, the OFRR inner surface was immobilized with streptavidin as the capture molecules. The sensorgram for the low biotin concentration (10 nM) was shown in Fig. 5 (b). It could be clearly seen that the biotin binding resulted in a 0.23 pm shift, corresponding to a surface mass density of 1.6 pg/mm2. The standard deviation, δ, was 0.0203 pm, leading to a surface mass density NEDL of 0.14 pg/mm2, which broke the 1 pg/mm2 barrier for many label-free optical biosensors [23].

In summary, we have experimentally demonstrated the excellent sensing capability of the ring resonator as a label free optical biosensor. These results set the benchmark for ring resonator sensing performance and compare very favorably with those obtained with other label-free optical sensors, which make the capillary-based optofluidic ring resonator a highly competitive sensing technology platform.

2.3 OFRR for chemical vapor detection

In the chemical vapor detection, the OFRR gas sensing platform is built by coating a vapor sensitive polymer layer on the OFRR capillary inner surface (as illustrated in Fig.6). The polymer-vapor interaction results in a change in the polymer thickness and refractive index, which in turn causes a spectral shift in the WGM that has the electric field present in the polymer layer.

Fig. 6. Schematic of an OFRR based chemical vapor sensor

OFRR based chemical vapor sensors exhibit several distinctive advantages compared to other optical vapor sensors [24]. 1) The OFRR vapor sensor is based on a glass capillary that bi-functionalizes as the gas microfluidic channel and the ring resonator sensor head as well. The sensor avoids the necessity of building extra bulky gas detection chambers commonly seen in chemical vapor sensors. 2) Rapid vapor detection is achievable owing to the efficient fluidics and the small cross-section size of the capillary, in which shorter diffusion time is needed for vapor molecules to reach the sensing polymer. 3) The OFRR enables on-column sensing capability. The vapor detection can be carried out at any position along the capillary using a tapered fiber to interrogate the vapor-polymer interaction. 4) The OFRR vapor sensor requires a sample volume as low as several microliters. 5) The OFRR vapor sensor is compatible with gas chromatography (GC) technique, which makes it a promising candidate for the development of μGC system. All of these characteristics are critical in the development of rapid, sensitive, selective, and portable vapor sensors.

We theoretically and experimentally investigated the performance of the OFRR-based chemical vapor sensors. Theoretical analysis shows the detection limit of the OFRR to be over ten times better than other ring-resonator technology [25], and experimentally the OFRR vapor sensor exhibited sub-second response time with a flow rate of only 1 mL/min and achieved a detection limit of 200 ppm in concentration or 1 ng in total mass for ethanol vapor [13].



Although the efficient fluidics and rapid detection makes the OFRR a promising platform for the development of vapor sensors, it suffers from limited selectivity in identification of the analyte, which is also a typical challenge for vapor sensors that rely on polymer as a sensing material. GC, on the other hand, is a well-established analytical tool that exhibits high selectivity by separating each analyte in the mixture and can be identified from their unique retention time. However, the traditional bench-top GC is bulky, power intensive, and requires long analysis time. It is desirable to develop the miniaturized, rapid, portable, and power efficient μGC analyzer. Since the OFRR achieves the dual-use of capillary as a ring resonator detector and the separation column, no extra micro detector is needed, so the system can be easily miniaturized. In contrast with other end-column detectors used in μGC system, OFRR based μGC is capable of performing multiple-point on-column detection in real-time. Therefore, the OFRR offers a unique solution to implement the μGC analyzer. Fig.7 shows the conceptual illustrations of a regular GC setup and the OFRR based μGC setup.

Fig. 7. (a) Conceptual illustration of a regular GC setup. (b) Cross-section view of a regular GC column. (c) Conceptual illustration of the OFRR-based µGC setup. (d) Cross-sectional view of the OFRR.

To demonstrate the proof-of-concept of the OFRR’s separation and on-column detection as μGC analyzer, several vapor analytes are used as a model system. In conventional GC and μGC systems, separation occurs as the sample travels through a separation column that is several meters to tens of meters long. In our experiment, the distance between the OFRR inlet and the detection fiber location is only ~2-3 cm. Fig. 8 shows separation of toluene/decane (Fig. 8(a)), hexane/decane (Fig. 8(b)), decane/dodecane (Fig. 8(c)), and toluene/DMMP (Fig. 8(d)). All the analytes could be well-separated at room temperature. The ability to separate several different samples over a short distance is enabled by the small diameter of the OFRR.

Figure 8. Separation and detection of different analytes with a PEG-400 coated OFRR μGC analyzer.

In the aim of developing a portable gas analyzer with high sensitivity, high specificity, and rapid response, we also studied the OFRR μGC analyzer for explosive detection [14]. Detection of explosives is a challenging task for the development of vapor sensors due to their extremely low volatility. Moreover, for practical applications, it is crucial for the sensor to discriminate the trace level of explosive substances out of other usually more volatile interferences in the air. DNT is chosen as a model system, because its chemical structure is similar to 2,4,6-trinitrotoluene (TNT) while its significantly higher volatility facilitated testing. Also, DNT is present in some explosives and is usually used as an

(a)

(d)(c)

(b)

(a) (b) (c) (d)



indicator to predict the existence of TNT. Fig. 9(a) shows the real-time response of the PEG 1000 coated OFRR sensor upon injection of DNT. A single sensing event is accomplished within only one minute at room temperature, showing great promise for detecting explosives in a rapid and simple manner. Concentration dependence on DNT mass is shown in Fig. 9(b), and a detection limit of approximately 200-300 pg for DNT vapor is achieved, which corresponds to about one second of solid-phase microextraction (SPME) extraction time from saturated headspace vapor at room temperature.

Figure 9. (a) Temporal response of the OFRR vapor sensor upon injection of DNT vapor. (b) OFRR vapor sensor

response to DNT vapor samples extracted with various SPME sampling times at room temperature. Reprinted from Ref. [14] with permission.

We further studied the capability of detecting DNT from a mixture of interferent compounds. Fig. 10(a) shows the separation between DNT vapor and nitrotoluene vapor. Although nitrotoluene and DNT only have one nitro group difference in structure, OFRR vapor sensors achieved very efficient separation between them. In Fig. 10(b), we further showed the highly reproducible and highly specific detection of DNT from a mixture containing three different analytes. The four traces not only demonstrated the separation capability among analytes that have relatively big difference in retention times, such as nitrotoluene and DNT, but also showed excellent separation efficiency in analytes that have similar interaction with the polymer, indicated by the well resolved peaks with one second retention time difference between nitrotoluene and TEP.

Fig. 10. Highly repeatable detection of DNT from interference background. (a) Separation and detection of DNT from

nitrotoluene/DNT mixtures. (b) Separation and detection of DNT from nitrotoluene/TEP/DNT mixtures. Inset shows the details of the first 15 seconds. Reprinted from Ref. [14] with permission.

The above OFRR based μGC achieved high performance for chemical vapor detection, however, it suffers from lack of resolving power for multiple analytes in a complicated mixtures due to shorter columns, which is an inherent constraint encountered in nearly all other μGC systems. In order to address this challenge, we developed a novel OFRR-based tandem-column μGC system to improve the separation resolution [15]. Via tapered optical fibers in contact with the OFRR, the WGM can be excited externally at any positions along the OFRR capillary, thus enabling multi-point, on-column, real-time detection of vapor molecules flowing through the OFRR. A schematic of the experimental setup of the

(a) (b)

(a) (b)



OFRR-based tandem-column µGC system is shown in Fig. 11. The first column (180 cm) is a conventional low polarity GC column, followed by a second, relatively short (~ 10 cm), home-fabricated OFRR column coated with a polar stationary phase. Two detection locations are chosen, one at the inlet of the OFRR, right after the first column, and the other one a few centimeters away from the first detection point to monitor the separation in the first column and the second column, respectively. Since OFRR on-column detectors record the retention time independently and provide complementary chromatograms for each chemical compound, co-elution at the terminal end of columns is no longer an issue for the OFRR based tandem-column µGC. By monitoring the retention time at multiple locations along the OFRR column, analytes can be well separated on at least one detection location.

Fig. 11. Conceptual illustration of tandem-column separation based on OFRR µGC

Fig. 12 shows a mixture of twelve compounds that are analyzed by the OFRR µGC system with tandem Rtx-1

column and PEG 1000 coated OFRR capillary column. All twelve compounds are well resolved after tandem-column separation, while two groups of co-elution appeared from the first column separation. Even without optimizing for speed, the analysis time for the twelve compounds is within 4 minutes, showing great potential of the tandem-column based OFRR µGC system for the development of a fast µGC.

Fig. 12. Chromatograms of twelve analytes with various volatilities and polarities obtained by the OFRR µGC analyzer with two detection channels. Reprint from Ref. [15] with permission In summary, we combined ring resonator sensing technology and GC separation technology and developed

OFRR-based µGC system. OFRR can either be used as a multi-dimensional on-column micro-detector to rapidly detect the separation from regular GC columns, or be independently worked as an integrated separation column and on-column detector to separate and detector the vapor mixture simultaneously. We have studied and demonstrated rapid detection of DNT at room temperature out of a volatile background using the OFRR µGC. Furthermore, we demonstrated rapid tandem-column separation and detection of vapor compounds using an OFRR-based µGC system, in which a nonpolar or slightly polar conventional GC column is used in conjunction with a short polar phase coated OFRR column so that the entire set of analytes can simultaneously be separated according to their respective vapor pressure and polarity in one



single analysis. A vapor mixture of twelve analytes of different volatilities and polarities are separated and detected within four minutes using the tandem-column OFRR-µGC system. All these work has significantly improved selectivity and chemical identification capabilities for chemical vapor mixtures. 2.4 Microfluidic lasers for intra-cavity biosensor

In the previous sections, we have developed OFRR-based label-free biosensors and OFRR-based µGC systems, where the OFRR is used as a passive micro-cavity. However, when the gain medium is introduced into the liquid core of the OFRR capillary (as illustrated in Fig. 13 (a)), the OFRR can be turned into an active microfluidic laser under optical pumping, which can be employed in bio/chemical sensing. Recently our lab has developed OFRR based microfluidic laser technology [16, 26]. It provides several distinctive advantages: 1) Similar to other optical ring resonators, the OFRR has extremely high Q factors (>107), thus providing an excellent feedback for low-threshold lasing (~ µJ/mm2). 2) It can operate with cavity in any liquid RI, and therefore be able to accommodate the low RI liquid such as water or buffer for biomolecular analysis. The reason that OFRR microlaser can uniquely operate with any liquid RI lies in the capillary wall thickness is sufficiently small, so that there is sufficient amount of light evanescently interacting with the gain medium for light amplification. 3) The laser emission can easily and highly efficiently be coupled out via an optical fiber, a fiber prism, or a waveguide in contact with the OFRR. Through such evanescent coupling, only lasing emission will be collected. The unwanted emission such as fluorescence background or ambient noise cannot enter the WGM and thus will not be transmitted through the optical fiber. As a result, the signal to noise is greatly increased. 4) Furthermore, the OFRR integrates naturally with the excellent fluidics inherent to capillaries, thus exhibiting superb fluid handling capability and allowing for approximately 1000 times reduction in sample effective detection volume (<10 nL). The unique characteristics of OFRR also enable to realize the microfluidic laser in a bio-amenable environment. We studied a bio-compatible optofluidic laser with dye molecule labeled single-stranded DNA (ssDNA) dissolved in water/buffer as the gain medium and with lasing through DNA-scaffold controlled fluorescence resonance energy transfer (FRET) [17]. Based on the bio-compatible optofluidic laser system developed, we further explored its capabilities for intra-cavity biosensing [18]. As a model system, the detection of DNA single-nucleotide polymorphism (SNP) with the OFRR laser sensor was demonstrated and exceptional detection selectivity was achieved as compared to the passive cavity sensing.

Fig. 13. Conceptual illustration of (a) OFRR based microfluidic laser setup. (b) Intra-cavity SNP detection with the OFRR laser. The OFRR cross-section forms the ring resonator. It supports the circulating optical resonant mode of high Q-factor, which interacts evanescently with the molecular beacon (MB) flowing through the OFRR and provides the optical feedback for MB to lase. Inset, MB fluorescence is quenched when it is in the closed state in which the fluorophore and quencher are in close proximity. MB fluorescence is restored upon hybridization with the target DNA. Blue arrow indicates the external laser pump.

Fig. 13 (b) illustrates the intra-cavity SNP detection scheme employing OFRR as the laser cavity. When placed in the OFRR, the molecular beacon (MB) becomes the gain medium of the OFRR laser, which is modulated by the MB conformational state determined by the hybridization with DNA molecules. Although small hybridization affinity difference between the target and single-base mismatched DNA causes only a small change in the laser gain coefficient, it is this small change that is nonlinearly amplified into orders of magnitude difference in the emission intensity due to

(a) (b)



the optical feedback provided by the laser cavity. Therefore, utilizing lasing emission as the sensing signal, the target and single-base mismatched DNA can be easily differentiated with the intra-cavity detection scheme. To demonstrate the SNP detection with the MB-OFRR laser, we first used simple DNA sequences of poly-A and its single-point mutation. The corresponding MB was labeled with FAM and DABCYL as the fluorophore and quencher. Fig. 14 (a) compares the MB emission in the presence of the target DNA and of the single-base mismatched DNA under otherwise the same experimental conditions. With the target DNA, multiple strong lasing peaks occurred at the longer wavelength side of the FAM emission spectrum, which is typical for a multi-mode dye laser like OFRR. In sharp contrast, only weak, featureless fluorescence was observed with the single base mismatched DNA (shown in Fig. 14 (b)). If we use the spectrally integrated emission intensity as the sensing signal, a discrimination ratio of 240 is obtained between the target and mismatched DNA, which represents more than two orders of magnitude improvement over the conventional fluorescence-based SNP detection using the same MB with a discrimination ratio of 2.

Fig. 14. Comparison of the MB (a) lasing emission spectrum and (b) fluorescence spectrum in the presence of the target DNA and of the single base mismatched DNA. The DNA-to-MB molar ratio is 1:1. MB concentration was fixed at 50 µM. Reprint from Ref. [18] with permission.

In addition to the MB emission intensity, the lasing threshold can be used to differentiate the target and single-base mismatched DNA. Fig. 15(a) shows the output intensities extracted from the MB emission spectra as a function of the pump energy density. For the target DNA, the lasing threshold of approximately 1 µJ/mm2 was achieved (see the inset of Fig. 15(a)). When the pump energy density was above 10 µJ/mm2, the laser output started to saturate. In contrast, no lasing emission was observed for the mismatched DNA, even with the highest possible energy density from the pump laser (100 µJ/mm2). Such huge difference in the lasing threshold is due to the inability of the mismatched DNA to open adequate MBs. With the increased DNA concentrations while keeping MB concentration unchanged, the laser can be achieved for both target and mismatched DNA, as shown in Fig. 15(b)-(d), with progressively decreased lasing threshold.

Fig. 15. Spectrally integrated MB emission intensity vs. the pump energy density for various concentrations of the target and single-base mismatched DNA. (a)-(d), [DNA]:[MB]=1:1, 2:1, 3:1, and 5:1. Reprint from Ref. [18] with permission.

(a) (b)



Direct and specific detection of the target DNA from a pool of single-base mismatched DNA molecules provides an excellent test-bed to examine the SNP detection selectivity. Unfortunately, to date, nearly all conventional MB SNP detection schemes can only detect the pure target or mismatched DNA individually and then theoretically estimate the fractional target DNA can be detected [27], since the signal from a DNA mixture becomes very difficult to de-convolve due to the small fluorescence difference between the target and the mismatched DNA. In contrast, as the intra-cavity detection provides the “clear-cut” analog-to-digital type conversion, the signal generated by the mismatched DNA is completely suppressed and thus does not interfere with the positive identification of the target DNA. To verify this, in Fig. 16, we set the pump intensity slightly higher than the threshold for the target DNA at 2:1 molar ratio with respect to the MB, but lower than that for the mismatched DNA. As a negative control, in Fig. 16(a)-(e) no lasing emission was observed for the single-base mismatched DNA even at an extremely high concentration ([Single-base mismatched DNA]:[MB]=100:1). However, strong lasing emission emerged when small amount of the target DNA was added, thus allowing us to positively identify the presence of 1 target DNA out of 50 single-base mismatched DNA molecules.

Fig. 16. Selective detection of PM DNA from high concentrations of SM DNA. a)–e) Negative controls. [PM DNA]/[SM

DNA]/[MB] is labeled on the right side of the corresponding graphs. f) Positive control. The MB concentration was fixed at 50 µM and the pump energy density was 0.8 µJ/mm2. Reprint from Ref. [18] with permission.

The OFRR laser can further be used to quantify the target DNA concentration, to detect SNP with more complicated DNA sequences, and even to function in biological buffer like serum [18]. To the best of our knowledge, this work marks the first application of active intra-cavity biomolecular detection using optofluidic lasers. It proves the tremendous capability of the optofluidic laser in biochemical analysis. The same detection scheme can readily be adapted for all different types of optofluidic lasers for highly sensitive and selective intra-cavity bio-analysis, thus opening a door to a number of research areas in optofluidics.

3. OPTOFLUIDIC MULTI-HOLE CAPILLARIES

A new flow-through multi-hole capillary has been recently developed in our lab for optofluidic applications [19, 20]. Compared to single-hole capillary with microfluidic channel, the multi-hole capillary consisting of a bundle of multiple flow-through micro/nano-sized holes provides a lot of unique features, including 3-dimensional (3-D) geometry for large surface-to-volume ratio, low sample volume (~ nL), inherent micro-/nano sized fluidic holes (or channels) for convenient and efficient sample delivery as well as high analyte capture efficiency, and light guiding properties for large signal accumulation. Combing with the state-of-the-art photonic technologies (such as Fabry-Pérot resonator, surface-enhanced Raman scattering, and fluorescence spectroscopy), we have developed the corresponding optofluidic devices. In the following, we will first describe the fabrication of the microstructured multi-hole capillary, and then explore its applications as an optofluidic FP based label-free biosensor, and a novel optofluidic SERS sensor.



3.1 Fabrication of microstructured multi-hole capillaries The microstructured multi-hole capillary was fabricated using an in-house computer controlled fiber/capillary drawing system (as shown in Fig. 3(a)) and a borosilicate glass preform. However, instead of using two CO2 lasers to heat only two sides of the preform (OD: 1~5 mm), we adopt electrical coil to heat the whole preform, where the preform is surrounded by the electrical coil and heated uniformly. In addition, to keep a certain air pressure inside the channels, a suitable length preform is used and sealed both ends using glue. Suitable heating time, feeding and pulling speeds are chosen to pull a long-length and large surface-to-volume ratio capillary. Fig. 17 (a) showed a pulled capillary that can be up to 50 cm long, and Fig. 17 (b)-(d) shows the SEM images of the pulled capillaries with 2700 holes (OD: 190 µm and ID: 2.9 µm), 7000 holes (OD: 200 µm and ID: 1.8 µm), and 300,000 holes (OD: 500 µm and ID: 0.7 µm), respectively. These results show that the micro/nanofluidic multi-hole capillary could be mass-produced easily and cost-effectively using the fiber drawing method with narrow hole size distribution.

Fig. 17 (a) picture of a pulled capillary. (b)-(d) SEM images of a 2700-hole capillary (OD: 190 µm and ID: 2.9 µm), a 7000-hole capillary (OD: 200 µm and ID: 1.8 µm), and a 300,000-hole capillary (OD: 500 µm and ID: 0.7 µm), respectively. The insets show the enlarged images of the holes.

3.2 Optofluidic Fabry-Pérot cavity sensor Effective integration between label-free sensors and micro-/nanofluidics is highly desirable for efficient sample

delivery to achieve rapid and sensitive detection. However, currently most of the label-free sensors employ the “flow-over” scheme, which relies on the analytes in bulk solution to diffuse to the sensing surface and suffers from mass transport problems. In order to address the inefficient analyte delivery issues, recently the “flow-through” scheme has been implemented in several nanoplasmonic sensors and photonic crystal nanofluidic sensors [28], where nanohole arrays are used as integrated nanofluidic channels for analytes to flow through.

Most recently we proposed and developed a new flow-through design that is micro/nanofluidic Fabry-Pérot (FP) cavity sensor using a microstructured multi-hole capillary (see examples in Fig. 17) [19]. As illustrated in Fig. 18 (a), the optofluidic FP sensor employs a microstructured multi-hole capillary with many built-in micro/nano-sized flow-through holes. When the capillary is placed between two reflectors, a FP cavity forms and detects the analytes binding to the internal surface of the holes. The optofluidic FP sensor has a number of advantages over the current label-free sensors. 1) As opposed to the planar surface detection in most of the sensors, micro/nanohole arrays enable 3-dimensional detection, which greatly enhances the detection sensitivity. While preserving the high sensitivity associated with the nanoporous silicon sensor, it significantly expedites the detection speed (minutes vs. hours) due to its built-in flow-through fluidics (rather than dead-ended holes in the nanoporous silicon sensor). 2) The micro/nanofluidic capillary can be mass-produced easily and cost-effectively from a preform using the fiber drawing method with much narrower hole size distribution than the nanoporous silicon sensor. 3) The hole size can be adjusted to accommodate a wide range of biological samples without significantly altering the setup or operation wavelength. 4) Finally, the Q-factor of the FP resonator can be as high as 27,000 [29], about three orders of magnitude higher than the aforementioned flow-through sensors, which, together with the higher sensitivity, should lead to much better sensing performance.

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sensor’s spectral shift as a function of the adsorbed layer thickness for different hole sizes. Holes are square and the wall thickness t is the same as the hole size d. d =1, 0.5, 0.2, 0.1, 0.05, 0.02 µm for Curves (1)-(6). λ0=1550 nm, ∆n= 0.2, L=100 µm. The inset illustrates the hole arrangement. Reprint from Ref. [19] with permission.

The surface detection sensitivity of the optofluidic FP sensor can be deduced by considering the mode spectral

shift caused by the fractional electro-magnetic energy change inside the cavity due to the biomolecular attachment:

0/ ( / ) ( / ) ,t A V n nδλ δ λ= ⋅ Δ ⋅ (2) where V and A are the total volume and internal surface area within the FP cavity, respectively. Eq. (2) shows that the surface detection sensitivity is linearly proportional to the surface-to-volume ratio, A/V. Similarly, the bulk refractive index sensitivity (assuming that the RI of all fluidic channels is homogeneously changed), nδλ δ , can be derived as:

0/ ( / ) ( / ),n W V nδλ δ λ= ⋅ (3) where W is the total volume of the fluidic channels within the FP cavity. The bulk RI sensitivity can be used to characterize the capability of detecting large species (such as virus).

Fig. 18 (b) plots the simulation results based on Eq. (2) for a FP cavity with square holes. In simulations, the length of the FP cavity is fixed to be 100 µm, and each hole has the same thickness as the solid wall (i.e. d = t). Therefore, when the hole size d is 1, 0.5, 0.2, 0.1, 0.05, 0.02 µm, the corresponding number of holes N within the resonance cavity is 50, 100, 250, 500, 1000 and 2500, respectively. With the decreased hole size, more holes and hence more surface area can be placed inside the cavity, resulting in a larger surface detection sensitivity. For the hole size from 1 µm to 20 nm, the sensitivity of 0.47 nm/nm to 23 nm/nm can be achieved. The sensitivity for the 20 - 100 nm holes is on par with the nanoporous silicon sensor with the similar pore sizes [30], but much larger than nearly all other types of label-free sensors such as surface plasmon resonance (SPR) based sensors, photonic crystal sensors, and ring resonator sensors [23].

To experimentally verify the feasibility and capability of the FP based flow-through sensor, we built a robust optical system as Fig. 19 (a) shows. A piece of 6-mm long micro/nanofluidic capillary (shown in Fig. 17(c)) was placed between two gold coated single mode fibers. The light from a tunable laser (1520 nm to 1570 nm) was coupled into one fiber and the transmitted light was collected by the other one. One end of the capillary was connected to a sample reservoir, and the sample was withdrawn quickly by vacuum from the other end with a flow rate of ~ 0.5 µL/min. Fig. 19 (b) shows the transmission spectra of the FP cavity in the absence and presence of the microstructured capillary (filled with DI-water). The presence of the capillary increased the cavity optical length, thus resulting in a smaller free spectral range. Note that even with thousands of holes inside the cavity, the FP resonance still persisted with a full-width-at-half-maximum of about 3.1 nm, corresponding to a Q-factor of 500 (much higher than other flow-through sensors).

We then characterized our sensor’s bulk RI detection capability by flowing a series of ethanol and DI-water mixtures. Fig. 20(a) shows that 550 nm/RIU (refractive index units) can be obtained, close to our theoretical estimation of 640 nm/RIU using Eq. (3). As shown in the inset of Fig. 20(a), it took less than 15 seconds for 5% ethanol to completely replace the DI-water initially filled inside the capillary. This represents the quickest analyte delivery rate among all flow-through nanohole sensors, as our sensor does not require any external microfluidic channels to be connected to the nanoholes.

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Fig. 19. (a) Pictures of the experimental setup. (b) Normalized transmission spectra of the FP cavity in the absence/presence

of the capillary. Reprint from Ref. [19] with permission.

To characterize the surface detection capability of our sensor, we measured the biomolecules attached to the surface. 1 mg/mL EZ-Link® Sulfo-NHS-LC-LC-Biotin (sulfosuccinimidyl-6-[biotinamido]-6-hexanamido hexanoate) (MW=670 Da) in PBS was flowed into the capillary that was functionalized with aminopropyltrimethoxy-silane (APTES). Fig. 20(b) shows that the wavelength shift increased as more and more analytes covalently bound to the sensing surface until a saturation shift of 0.35 nm was reached in about 7 minutes. Assuming that a fully packed analyte layer was formed on the sensing surface with the thickness increase corresponding to the length of analyte (i.e., 3.05 nm) and that the analyte has the same RI as silica, we arrived at the surface detection sensitivity of 0.115 nm/nm, close to 0.124 nm/nm based on Eq. (2) discussed earlier.

Fig. 20. (a) Characterization of the sensor’s bulk RI sensitivity. Inset shows the sensorgram for 5% ethanol flowing through the capillary to replace DI-water. (b) Sensorgram of 1 mg/mL Sulfo-NHS-LC-LC-Biotin binding to the silanized silica sensing surfaces inside the capillary. Reprint from Ref. [19] with permission.

In summary, we have presented a robust optofluidic FP cavity sensor with integrated flow-through multi-hole capillary, which exhibits rapid analyte delivery (< 15 s), high bulk RI sensitivity (550 nm/RIU), and potentially high surface detection sensitivity (> 10 nm/nm). The results show that this sensor provides a very promising platform for rapid, sensitive, and high-throughput biological and chemical sensing.

3.3 Optofluidic SERS based biosensor

While significant progress has been made in 3-D optofluidic SERS systems, they still experience either relatively low limited sensitivity, complicated and costly device fabrication procedures, short interaction length unfavorable for large signal accumulation or in-line analyte monitoring, difficulties in controlling metallic nanoparticle aggregations and

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nanoclusters, or time consuming SERS substrate (i.e., metallic nanoparticle) deposition and analyte accumulation processes. How to achieve a robust, simple, reliable, highly sensitive, and cost-effective optofluidic SERS platform still remains unanswered.

With the unique flow-through multi-hole capillary, we proposed and developed a novel optofluidic SERS sensing platform for rapid, reliable, and ultrasensitive analyte (in liquid/vapor phase) detections [20]. It provides a number of distinctive advantages compared to the state-of-the-art technologies: 1) The capillary contains thousands of micro-/nano sized holes, providing large surface SERS-active sites to achieve much higher SERS intensity than the widely-studied planar SERS substrates. 2) The inherent flow-through micro/nano-sized holes (or channels) enable robust and reliable metallic nanoparticle/nanocluster immobilization, and fast and convenient sample delivery for subsequent in-line SERS detection, a big advantage over other planar SERS substrates [31] and dead-ended 3-D porous structures [32]. 3) Additionally, the micrometer- and nanometer-sized holes combines convective flow and short diffusion length scales, which significantly reduces both the time required for analyte molecules to reach the SERS-active surfaces and the sample consumption [28]. 4) The excitation light and Raman-scattered photons can propagate along the entire length of the capillary [33], which greatly increases the SERS accumulative length as the wide-studied photonic crystal fiber (PCF) [34] does but with much larger sensing surface area due to more available SERS-active channels. 5) The multi-hole capillary does not need delicate effort to arrange the holes to construct the photonic crystal (PC) structure as PCF does, and can be massively fabricated with the fiber drawing method as for PCF, but more easily and more cost-effectively. Fig. 21 shows the schematic diagram of a Raman detection system together with depiction of the light process in the multi-hole capillary in two detection schemes: transverse and longitudinal directions, respectively. Both of the detection configurations provide unique advantages. In the transverse detection, the SERS signal is from the SERS-active sites within the detection volume determined by the laser excitation and the SERS collection optics. Due to the extremely large surface resulting from thousands of micro/nano-sized holes in the capillary, 3-D sensitive SERS detection can be achieved. The longitudinal detection takes additional advantage of SERS signal accumulation along the capillary.

Fig. 21. Schematic of a flow-through optofluidic Raman system based on a multihole capillary. The SERS excitation and detection are arranged or transversely (top) or longitudinally (bottom) with respect to the capillary. Reprint from Ref. [20] with permission.

In our experiments, Raman and SERS measurements were carried out with the customized Raman spectroscopy system, consisting of a 785 nm excitation laser and a spectrometer. An aspheric lens was used for delivery of the laser excitation and collection of the Raman signal. The excitation light was focused onto a spot of approximately 3 µm in diameter. All the Raman data was acquired with 6 mW excitation power and 2 seconds exposure time. The flow-through multihole capillaries used in our experiments were 190 µm in outer diameter and had 2700 uniform 2.9-µm holes with 0.7-µm thick wall between two adjacent holes (see Fig. 17(a)). To realize a SERS-active capillary, we used polyelectrolyte-mediated surface immobilization of noble nanoparticles onto the capillary surface [35]. Poly(allylamine hydrochloride) (PAH) was first allowed to adsorb at the surface of the flow channels, served as anchoring site for Au nanoparticles, and then 2 µL Au nanoparticle solution was flowed through the capillary within 5 minutes. Fig. 22(a) and



(b) show the typical immobilization condition for Au nanoparticle solution with a concentration of 2.4×109 particles/mL. Almost no Au nanoparticle was observed on the capillary facet as there was no PAH modification, while inside the flow channels, discrete Au nanoparticles were uniformly covered on the surface area with a density of 6.1+0.5 particles/µm2, except a few sparked dimers or trimers.

Fig. 22 (a) The facet of the capillary adsorbed with gold nanoparticles. (b) Cross section of flow channels with adsorbed gold

nanoparticles. Reprint from Ref. [20] with permission.

We first investigated the proposed SERS platform using the transverse detection. The effective detection volume is approximately 850 µm3 (i.e., 0.85 pL) which is determined by the excitation laser spot size and Raman signal collection optics (120 µm collection depth across the capillary). Since the multihole capillary provides a very large surface-to-volume ratio of 0.87 µm-1, the total effective illuminated surface area is estimated to be 740 µm2, over two orders of magnitude larger than 7 µm2 obtained with a single flat surface using the same excitation and collection optics. The performance of the flow-through multihole capillary for Raman spectroscopy was then characterized with the well-studied molecular probe R6G dye. Typical SERS spectra for R6G at the concentration of 5×10-7 M were taken stepwise at a 160 µm interval from the capillary head to a position 1.6 mm downstream, as plotted in Fig. 23(a). Since nanoparticle transport is dominated by convection in the microfluidic channel, the gold nanoparticle density adsorbed onto the PAH-modified channel surface gradually decreases along the flowing direction, as reflected by the gradually reduced SERS signal in Fig. 23(a). Note the decrease in the SERS intensity is monotonic without any abrupt change, suggesting that the SERS signal results mainly from discrete single nanoparticles rather than nanoclusters or aggregations, which would cause large fluctuations in the SERS intensity distribution along the capillary.

Fig. 23 (a) Series of SERS spectra for 5×10-7 M R6G obtained by the stepwise measurement along the capillary in an increment of

160 µm. (b) SERS spectra of different concentrations of R6G (0, 5×10-13 M, 5×10-12 M) using SERS-active capillary, and normal Raman spectrum (NRS) of 10-2 M R6G in solution in a 1-mL glass cuvette (with 10 times magnification). Reprint from Ref. [20] with permission.

To investigate the capability of the flow-through optofluidic SERS platform, we measured a series of low concentrations of R6G solutions from 5×10-10 M to 5×10-13 M. Fig. 23(b) shows the SERS spectra of R6G at concentrations of 5×10-12 M and 5×10-13 M. For comparison, the SERS spectrum of pure gold nanoparticles in the absence of R6G and the normal Raman spectrum (NRS) of R6G at a concentration of 1×10-2 M in a 1-mL glass cuvette

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are also shown. The characteristic R6G peaks are clearly observed for 5×10-12 M, despite some differences from the normal Raman spectrum, which are expected as the interaction between the molecules and gold nanoparticles results in conformation changes and variations in molecular vibration modes. For R6G with the concentration of 5×10-13 M, the SERS peaks at 785 cm-1 and 1369 cm-1 are still observable to indicate the presence of R6G molecules, attesting to the excellent sensing performance of the multi-hole capillary as the optofluidic SERS platform.

The enhancement factor (EF) of the proposed SERS system can be estimated using the following equation: (4)

where ISERS and Inorm are the measured Raman intensity for SERS and NRS, respectively. Nnorm and NSERS are the number of R6G molecules in the detection volume for NRS and SERS, respectively. We used the strongest signature stretching mode at 1369 cm-1 at 5×10-12 M for SERS and 1×10-2 M for NRS shown in Fig. 3(b). The number of R6G molecules detected is 5.6×103 for 5×10-12 M solution in SERS and 5.1×109 for 1×10-2 M solution in NRS. With 35 times higher SERS intensity than NRS intensity, the EF is calculated to be 3.2×107. We estimate that over 106 fold enhancement results from the gold nanoparticle adsorbed on the capillary surface [35]. Additional enhancement may be attributable to the light confinement and multiple surface reflections or scattering of the multihole capillary. To verify this, a control experiment was performed to measure the normal Raman spectra of 1×10-2 M R6G in the multihole capillary, which showed over 20 fold enhancement in NRS could be achieved as compared to the same experiment performed using a 1-mL glass cuvette (see Fig. 24a).

Fig. 24 (a) The normal Raman spectra of 10-2 M R6G loaded in a 1-mL glass chamber and the multi-hole capillary, respectively. (b)

SERS intensity at 1369 cm-1 for a series of concentrations of R6G from 5×10-13 M to 5×10-10 M. Error bars were obtained with at least 3 runs. The solid curve shows the Langmuir isotherm fit to the experimental data. Reprint from Ref. [20] with permission.

The limit of detection (LOD) for the multihole capillary optofluidic SERS system was also evaluated by varying R6G concentrations, then measuring the intensity of characteristic R6G band at 1369 cm-1 at a position with maximum signal along the capillary (mostly close to the flow inlet), as plotted in Fig. 24(b). The experimental data was well fit by a Langmuir isotherm curve with nonlinear least-squares regression. By substituting 3 standard deviations of the blank signal into the fitted Langmuir isotherm curve, an estimated LOD of 70 fM is obtained, which is three orders of magnitude better than that achieved with a long PCF with discrete silver nanoparticles [36, 37] and similar to that in the nanoporous polymer monolith with silver nanoparticle aggregations [38].

Apart from excellent performance achieved with the transverse direction, further improvement is possible with the longitudinal measurement by accumulating SERS intensity along the capillary. We used a 3-mm-long multihole capillary loaded with a 2 µL mixture with the final concentration of 5×10-12 M R6G and 2.4×109 particles/mL gold nanoparticles. We first measured the SERS intensity at several positions along the capillary using the transverse detection and found continuous SERS signal along the capillary (Fig. 25(a)). Then we switched to measure the SERS intensity along the longitudinal direction. Even without optimization of the capillary (such as endface flatness) for better SERS collection efficiency, more than 5.5 times higher SERS signal was achieved than the maximum signal in the transverse direction (Fig. 25(b)), which suggests that the capillary has the ability to accumulate SERS signal along the capillary length. The total EF in the longitudinal direction is estimated to be over 108, and the LOD is estimated to be 14 fM.

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In summary, we have developed a novel optofluidic SERS platform using a flow-through multi-hole capillary. It also allows for cheap, fast, controlled, reproducible, easily fabrication, and low sample consumption. Moreover, ultrasensitive SERS detection with an enhancement factor of over 108 and a detection limit of less than 100 fM for R6G molecules, has been achieved. In addition, with the developed optofluidic SERS sensor, we also successful achieved rapid (within 1 second of exposure) and ultrasensitive (at a sub-ppb level) vapor detection in open air at room temperature. We believe that the robust SERS-active optofluidic flow-through multi-hole capillary has the potential to dramatically increase the applicability of SERS-based chemical and biological detection and enable rapid and low cost sample analysis.

Fig. 25 (a) SERS intensity distribution along the multihole capillary in the transverse direction. (b) Comparison of SERS intensity in

the transverse and longitudinal detection using the same multihole capillary adsorbed with the mixture of R6G molecules and gold nanoparticles. Reprint from Ref. [20] with permission.

4. CONCLUSIONS

In conclusion, we have successfully developed a number of optofluidic devices with the unique capillary structures (thin-walled single-hole capillary and multi-hole capillary). These capillary structures not only provide superior fluidic handing capability, but also enable to combine with the state-of-the-art photonic technologies (including ring resonators, FP resonator, SERS, etc.), which make them suitable for detecting various biological and chemical analytes in either liquid or vapor phase. Moreover, the optofluidic devices are very robust, and provide facile integration of the sensing elements into a range of capillary and microfluidic devices for biomolecular detection. For example, integrating with micro-gas chromatography (µGC), we developed optofluidic ring resonator based on-column µGC analyzer. Furthermore, due to its small size and simple configuration, it is also promising to form arrays of sensors to provide high-throughput detections. We believe that these optofluidic platforms we developed have the potential to dramatically increase the applicability of optofluidics-based chemical and biological detection and enable rapid and low cost sample analysis.

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