We demonstrate a novel integrated optical fiber interferometer for in-fiber optofluidic detection. It is composed of a specially designed hollow optical fiber with a micro-channel and two cores. One core on the inner surface of the micro-channel is served as sensing arm and the other core in the annular cladding is served as reference arm. Fusion-and-tapering method is employed to couple light from a single mode fiber to the hollow optical fiber in this device. Sampling is realized by side opening a microhole on the surface of the hollow optical fiber. Under differential pressure between the end of the hollow fiber and the microhole, the liquids can form steady microflows in the micro-channel. Simultaneously, the interference spectrum of the interferometer device shifts with the variation of the concentration of the microfluid in the channel. The optofluidic in-fiber interferometer has a sensitivity of refractive index around 2508 nm/RIU for NaCl. For medicine concentration detection, its sensitivity is 0.076 nm/mmolL−1 for ascorbic acid. Significantly, this work presents a compact microfluidic in-fiber interferometer with a micro-channel which can be integrated with chip devices without spatial optical coupling and without complex manufacturing procedure of the waveguide on the chips.
© 2017 Optical Society of America
The integration of microfluidics and optics has led to a new interdisciplinary of optofluidics. Such optofluidic integration of photonic functions with non-solid media has made tremendous progress in recent years [1–4]. The advantages of microfluidics and optical waveguides are integrated in a common device to realize the interaction between microfluids and light with high specific surface area [5, 6]. Within optofluidic devices, high sensitivity and high spatial resolution based on the structures such as resonators, gratings, and photonic crystals can be obtained based on optical analysis methods. In them, the light propagation is controlled by the distribution of the refractive indices (RIs) of the waveguide and affected in the form of reflection, transmission, or attenuation by the interaction between the light and the analyte [7–9].
On the other hand, as a kind of special waveguide, microstructured optical fibers (MOFs) have micro-channels around the cores between the two ends and hence aroused significant interest in chemical and biological sensing technologies [10–13]. This kind of fibers are favourable candidates for optofluidic devices and well suited for microanalysis. For example, some of them are based on solid-core photonic crystal fibers (SC-PCFs) which evanescent field of the cores are sensitive to the material within the holes [14, 15]. In these structures, light is guided by total internal reflection (TIR). In addition, in liquid-filled hollow-core photonic crystal fiber, light guidance can be either achieved in the hollow-core due to the photonic bandgap  or anti-resonant-reflection effect , the light can be guided by TIR in the junctions of the microstructure  or by TIR in the hollow-core if it is selectively filled with a liquid . Specific benefits which cannot be found in standard optical fibers are present by them. Including as waveguides, the MOFs also provide long-path cells in the diameter of micrometers to enhance the evanescent field effect which offer the direct interaction between the light and analytes during samples analysis [19–21]. Specifically, compared with macrosystems, devices based on MOFs only consume trace reagents and is conducive to high efficient mass transfer and real-time on-line detection.
Furthermore, optical fiber interferometers are deeply investigated in the detection of physical parameters, chemical concentrations, and biological molecular [22–27]. In comparison with the spatial interference light path, the optical fiber has the advantage to reduce the wavefront distortion of the atmospheric turbulence. They have a fast response to internal or external environment and they are immune to electromagnetic noise disturbance. Because the phase-detection is independent of the intensity of the light, the device based on fiber-optic interferometer will present excellent stability and the signal will not be influenced by the instability of the light path [28, 29]. Among them, twin-core fibers are attractive as the components of optical fiber interferometers because of the integration of the optical path [25, 30]. However, most of the current fiber interferometers used as sensors have open structures and need extra flowing cells.
In this paper, an in-fiber interferometer integrated optofluidic device based on a specially designed hollow optical fiber with two cores is proposed. In-fiber optofluidic detection based on phase detection of the spectrum is first realized in this microfluidic device. It enables an adequate light coupling of the fluid and waveguide through the evanescent field around the core in the micro-channel of the optical fiber. When the microfluid passing through the micro-channel, the effective RI around the core can be modulated. Then, by changing the optical path length difference (OPD) between the core in the micro-channel and the core in the annular cladding, interference spectrum shifts and reflects the information of the microfluid in the optical fiber. Moreover, this in-fiber interferometer with a compact size can be easily connected and integrated with the microfluidic chips following the unit of the microreactor, micromixer, or microseparator to finish the detection. The optofluidic device has great potentials in the analysis of chemistry, drug, biology and environment.
The in-fiber microfluidic interference device is based on a novel hollow twin-core optical fiber (HTCF) which is newly fabricated in our lab. The preform for it is fabricated by the assembling method following the usual way in our lab [31, 32]. By selecting suitable sets of parameters, such as heating temperature, translation speed, drawing speed and pressure in the hole, the fibers with the hole diameters around 40-70 µm can be obtained with the drawing furnace at a high temperature. The cross-section view of the optical fiber used in the experiment is shown in Fig. 1(a). This optical fiber has an air hole with the diameter of 42 µm and an annular cladding with the thickness of 41.5 µm. The diameters of the two cores and the cladding are 7.9 µm and 125 µm. The inner core in the air hole is far from the other core and they leave enough distance to avoid light coupling between each other. Importantly, one core which located in the annular cladding is served as a reference arm and another core which located on the inner surface of the micro-channel is served as sensing arm in order to be adjusted by the microfluid. 3D image of the RI for the optical fiber is shown in Fig. 1(b). The RI distribution of the sample fiber was measured by an RI profiler. One dimensional RI distribution profile along the diameter which passes through the two cores is shown in Fig. 1(c). We can observe the two cores (labeled as core 1 and core 2) have higher RI than the annular cladding. The RI of the cores is respective 1.462 and that of the cladding is 1.457.
The proposed optofluidic interferometer is schematically illustrated in Fig. 2(a). As shown in the figure, the light path of Michelson interferometer is composed of a supercontinuum laser source (Anyang laser SC-5, 1W), a 1x2 3 dB single-mode fiber coupler, a piece of HTCF and an optical spectrum analyser (OSA, AQ6370 C). One end of the HTCF is spliced to the fiber coupler and the melting point is drawn to a tapered region. The other end is covered with a layer of Au film by ion sputtering method to get a reflection at the silica-air interface. The device is immobilized on the silica substrate by adhesive.
Figure 2(b) presents a sketch of the coupling between single mode optical fiber and the HTCF. Broadband light from the supercontinuum laser is efficiently coupled into the two cores of the HTCF through the tapered region, and the reflected beam from the end surface of the HTCF is recombined when they transmitted through the taper again. Then, the interference spectrum of the interferometer is displayed by the OSA. Here, the taper plays a role of light beam coupler that split the light from the 3 dB coupler into the two cores of HTCF and recouples the two beams to form interference spectrum.
To build the microfluid path, the side wall of the HTCF is opened by etching a microhole on the surface of HTCF. Figure 2(c) shows the sketch of the setup. Specifically, to find correct direction of the HTCF for opening a hole without damaging to the cores, the end of HCTF was observed by a flattened microscope with a CCD camera. Before etching, the plane of the two cores is adjusted to parallel to the horizontal plane. Then, during the process of etching, the two cores are placed away from the processing point, so the structure of the waveguide is not damaged. Here, CO2 laser power is set to be 8 W with a speed of scanning of 100 mm/s. The frequency of the laser is 20 kHz. Then, the HTCF is placed at the focus point of the CO2 laser beam and is scanned in the direction of perpendicular to the fiber. The distance between the hole and the spliced point is about 2 mm. The etching process is observed under a microscope with CCD camera. The microhole can be obtained by repeatedly scanning for 120 times. As the inset of Fig. 2(a) shows, the width of the hole is about 50 µm and the morphology is uniform without obvious melting effect beside the etching point.
Sampling is realized under negative pressure with a vacuum pump through a PTFE capillary tube, which is connected with the microhole on the surface of the HTCF using a silica microtube. The end of the silica microtube is encapsulated using UV epoxy glue.
The fusion-and-tapering method was employed to couple light from SMF into HTCF . To realize the light distribution in the both cores of HTCF, a short piece of HTCF is melted and collapsed to the solid structure by a fusion splicer previous tapering. Then, the solid part of the HTCF is cut to leak the collapsed end face and spliced with the common terminal of the 3 dB coupler. The splice point of the fiber was heated and gradually drawn to form a tapered region. The light power was gradually transferred from the single-core single-mode fiber into the cladding mode near the waist region, and the optical power was gradually concentrated from the cladding mode (near the waist) into the higher refractive index zone of the two cores after the waist region. During tapering, the output powers of the two cores were being monitored with an IR camera (Electrophysics 7290A) and a beam view analyzer (Coherent Inc.). In this experiment, when the final diameter of the tapered fiber waist is about 30 µm, an optimal light splitting ratio of the two cores (about 50:50) was obtained.
The optical coupling between the core of SMF (belong to 3dB coupler) and the two cores of HTCF is simulated by the beam propagation method [34, 35] and shown in Fig. 3. Here, the diameter of the cores (including SMF and HTCF) is set as 7.8 µm and their RI is set as 1.462. The cladding around the cores including the collapsed region of HTCF is set as 1.457. The diameter of the taper waist is set as 30 µm and the diameter of the cores is reduced to 1.7 µm. In this tapered region, it can be observed from the left picture that most of the guided light inside the single core of the SMF is transferred into the two cores of the HTCF and is separated as two interference arms. The right plot shows the power distribution inside the three cores along the direction of propagation. The power at wavelength 1520 nm in the core of SMF gradually reduced with the diameter of the core decreases. Then, the power in each core of HTCF gradually increased to about 0.4 unit when they connected with the SMF at the waist region. Considering the mechanical strength of the device, the final diameter of the taper is not too small. When the taper is finer, the coupling efficiency will be further increased. Experiment results show that this diameter is enough to obtain obvious interference spectrum.
In this in-fiber integrated optofluidic interferometer, the output intensity of the interferometer can be expressed as:
where and are the light intensities in core 1 and core 2 of the HTCF, respectively. Δ is the phase difference between the two cores after propagating through the HTCF with a length of . It can be described as Eq. (2). Where is the central wavelength of the interference spectrum and and are the effective RI of the core 1 and core 2, respectively, is the RI difference between the two cores. From Eq. (2), a transmission dip corresponds to a phase difference of , and the wavelength of the dip can be described as Eq. (3). It can be seen that changes of or will bring a wavelength shift of . In this design, the length of the optical fiber will remain constant. However, will be changed when the microfluid samples are injected into the optical fiber and cause the response of interference spectrum.
3. Experimental results and discussion
The typical spectrum of the as prepared optofluidic device the in-fiber integrated optical fiber interferometer was obtained and shown in Fig. 4(a). We can observe the spectrum is uniform. In addition, several HTCFs with different length were tried in the device. In Fig. 4(b), it is obvious that the free spectrum range (FSR) changes with the length of the HTCF in the optofluidic device. We can observe the FSR decreases with the length of the part of HTCF in the device. When the length of the HTCF is 7.5 cm, the FSR is 9.8 nm. However, when the length of the HTCF is increased to 15 cm, the FSR decreases to about 5 nm. Considering trying to keep a longer path of the interaction between the sample and the interfering arm, a lower microfluidic resistance, and especially a bigger FSR for tracking the changes of the dip in the spectrum when testing, 10 cm of the HTCF (FSR is 7.5 nm) was finally chosen in this experiment.
In this in-fiber integrated optofluidic interference device, the fluid of sample was inhaled into the optical fiber through the microhole on the surface of the fiber. To demonstrate the Michelson interferometer’s RI sensitivity, a series of NaCl solutions with different RI were inhaled into the optical fiber. The microhole on the surface of the fiber was applied a vacuum pump with a pressure of 6x10−2 Pa. Then, the microfluid under the pressure difference between the microhole and the end of the HTCF was formed. When the sample solutions with different concentrations entered into the micro-channel of the fiber, the RI around core 2 displayed different values. That means it would cause phase shifts of the interference spectra. According to Eq. (3), the concentration of the samples could be reflected by the wavelength shift of . In the experiment, a NaCl solution with higher concentration in the sample chamber was continuously diluted by water to obtain solutions with different concentrations. Generally, the spectrum immediately shifted when the solution was diluted and was stable after 20 s. This means there was a mass transfer process of the microfluid in the fiber. Specifically, by diluting the solution with the concentration of 5% (w/w), a series of solutions with the RI range from 1.3421 to 1.3373 were obtained (measured by Abbe refractometer) and the corresponding dip of the interference spectra around 1520 nm was recorded at room temperature.
The change of the interference spectrum within the range of one FSR with different RIs of NaCl microfluidic solutions (labeled as sample A, B and C) is present in Fig. 5. From the curve, we can observe the distance between two dips is about 7.5 nm and spectrum present obvious shifts. It can be clearly observed the dip shifted towards shorter wavelength when RI was decreased from sample A to sample C. Notably, when the RI only changed 0.0018 from 1.3421 to 1.3403, the corresponding dip presented a big shift from 1520.5 nm to 1515.9 nm. By continuously tracking the position of the dip around 1520 nm when the RI of the microfluid decreased, the relationship between RI and the corresponding wavelength shifts were obtained and shown in Fig. 6 (error bars indicate the standard deviation of 5 measurements). When RI decreased from 1.3421 to 1.3373, the dip shifted from 1520.5 nm to 1508.6 nm. From the fitting line, this optofluidic device provides the sensitivity of 2508 nm/RIU (with the limit of detection of 2.2x10-6 RIU when the OSA working with a resolution of 0.05 nm). Comparing with the interferometer based on twin-core fiber without microhole , this sensitivity is greatly higher than that of 270 nm/RIU which is reported. This is due to the bare core in the microhole which provides enough area for contacting with the analyte.
To further demonstrate the application for quantitative detection of the drug, the microfluidic NaCl samples were changed into the specific ascorbic acid samples. The commercial ascorbic acid pills were washed under ultrasonic wave to get rid of the pigment and the sugar-coating on the surface. After drying under 50 for 2 h, a solution with the concentration of 200 mmolL−1 was prepared. After that, the solution was also continuously diluted with deionized water and the spectrum was continuously recorded. Similarly, the blue shift of the spectrum could be obviously observed during diluting the samples because of the changes of the RI of the samples. The variation of wavelength shift with different concentrations of ascorbic acid is shown in Fig. 7 (error bars indicate the standard deviation of 5 measurements). As the ascorbic acid sample was diluted from 200 mmolL−1 to 50 mmolL−1, the RI present a decrease of −0.0047 from 1.3384 to 1.3337. The corresponding dip of the spectra presents an obvious shift of −11.58 nm towards shorter wavelength. In the range of 50-200 mmol/L, this in-fiber optofluidic interferometer sensor shows well linearity with a slope (concentration sensitivity) of 76 pm/mmolL−1 with a minimum detectable concentration change of ~0.65 mmolL−1when the OSA working with a resolution of 0.05 nm.
In fact, the temperature crosstalk of an interferometer sensor is a key factor which determines the usability of the device. To investigate the temperature response of the in-fiber optofluidic interferometer, the whole device was placed in a thermostatic drying oven and gradually increasing the temperature from room temperature of 30 to 50 . As a result, in this temperature range, the transmission spectrum exhibited small shifts toward shorter wavelength which was observed when the temperature was increased, as shown in the inset of Fig. 8. From the fitting line, we can observe the dip presents a linear shift of −0.95 nm with the temperature sensitivity of 45 pm/ (the limit of detection of 1.1 when the OSA working with a resolution of 0.05 nm). Generally, the temperature-induced shift of the wavelength relays on the thermo-optic effect of the fiber cores [25, 30] and the thermal expansion of the interferometer. In this design, the two cores are located in the same fiber, and the thermal expansion for the two cores should be approximately the same. Then, maybe the cross-sensitivity to temperature was caused by the thermal-optic effect of the fiber core. If we define the cross-sensitivity of temperature to concentration of ascorbic acid is the result of temperature sensitivity (that is 45 pm/) divided by concentration sensitivity (that is 76 pm/mmolL−1) of ascorbic acid, and there is no temperature compensation in the above ascorbic acid testing, the concentration measurement cross-sensitivity is about 0.59 mmolL−1/. It means that the change in wavelength caused by per Celsius degree corresponds to the wavelength change caused by the concentration change of 0.59 mmolL−1. In other words, per Celsius degree will cause an error of 0.59 mmolL−1. Considering the limit of detection of ~0.65 mmolL−1, we believe the influence of the temperature below 1.1 (0.59 divided by 0.65) can be ignored.
The structure of this special optical fiber first simultaneously provides microfluidic channel and the waveguides to realize online optofluidic sensing basing on in-fiber interference which is hard to be realized in traditional optical fibers and other hollow fibers. Especially, the two cores of the optical fiber perform as two interference arms and the micro-channel provide a long sensing region for the microfluid. The liquids can form steady microflows in the full length of the optical fiber interferometer under a differential pressure between the inlet and the outlet on the surface and the end of the optical fiber, and can sufficiently contact with the sensing arm in the micro-channel. The results show that the interference spectrum of the in-fiber optofluidic interferometer shifts with the variation of the concentration of the analytes in the HTCF. The device has good operation linearity and exhibits a high RI sensitivity of 2508 nm/RIU. Further, quantitative detection of ascorbic acid has also been demonstrated in this optic interferometer sensor with the sensitivity of about 0.076 nm/mmolL−1. In addition, the optofluidic in-fiber interferometer presents a low temperature cross-sensitivity of 0.59 mmolL−1/ for the detection of ascorbic acid. For the qualitative detection of the samples, molecular probes can be further immobilized in the HTCF to realize selectivity. for drugs and biomolecules. In comparing with the interferometer based on twin-core fiber without microhole , this device presents higher sensitivity. Although the sensitivity is not as high as the reported ultra-sensitive interferometer (10981 nm/RIU) based on a twin-core fiber which is side opened femtosecond laser , such a low cost device realizes microfluidic in-fiber detection with a micro-channel based on interference spectrum and has high potentials in integration with microfluidic chips for chemical and biological analyse.
National Natural Science Foundation of China (NSFC, 61405043, 11574061, 61290314, 61535004, 61307005, 61227013, U1231201); China Postdoctoral Science Foundation (2014M551217); the Natural Science Foundation of Heilongjiang Province (F201405) and the Fundamental Research Funds for the Central Universities (GK2110260186).
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