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Abstract
Temperature cross-sensitivity is a long-standing challenge for most of the in-line fiber optofluidic waveguide biosensors. In this paper, we propose a dual-optofluidic waveguide antiresonant reflecting optical waveguide (ARROW) biosensor for the detection of interferon-gamma (IFN-γ) concentration with temperature compensation. Two Fabry-Perot resonators infiltrated with IFN-γ and NaCl were formed in a hollow core fiber, which generate two resonance dips based on the ARROW model. The optical biosensor for the detection of interferon-gamma (IFN-γ) has been a key research interest in recent years because IFN-γ is an important early biomarker for many serious human diseases. Based on the dual-optofluidic waveguide ARROW biosensor, the IFN-γ concentration can be measured through the modulation of the resonance condition of the ARROW, while the temperature fluctuation can be eliminated due to same thermo-optic coefficients of two infiltration liquids. The experimental results show that the response of the ARROW biosensor can be amplified significantly with the signal-enhanced streptavidin, and the limit of detection of 0.5 ng/ml can be achieved for the IFN-γ concentration. More importantly, the influence of the temperature could be compensated through the referenced resonance dip. The proposed fiber biosensor has a great potential for the real-time detection of IFN-γ concentrations in the fields of health monitoring, cancer prevention, biological engineering, etc.
© 2020 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
1. Introduction
In-line fiber optofluidic waveguide biosensors possess both the enhanced sensing performance and ultracompact size, which have been widely used in the lab-in-fibers chemical and biological sensing. The in-line fiber optofluidic waveguide biosensors combine the microfluidic channel and light waveguide in signal optical fiber, which have been widely used in the lab-in-fibers chemical and biological sensing [1]. The optofluidic waveguides are formed by microstructured holes in the fiber. Compared with conventional on-chip optical biosensors, in-line fiber optofluidic waveguide biosensors not only simplify the sensing configuration with small size and easy operation, but also improve the sensor sensitivity significantly due to the enhanced interaction between the guide light and biomolecules. In-line fiber biosensors possess many distinctive advantages, such as high sensitivity, compact size, and immunity to electromagnetic interference [2].
Interferon-gamma (IFN-γ) is a typical cytokine that is secreted by T cells, natural killer cells, and mucosal epithelial cells [3]. Due to the significant effect on immunostimulatory and immunomodulatory, IFN-γ has been regarded as an important early biomarker for many serious diseases, such as human immunodeficiency virus (HIV) [4], pleural effusion of tuberculosis [5], and inflammation [6], etc. For instance, a concentration of 1 ng /mL for the IFN-γ indicates the potential of Johne's disease for the patient [4]. Therefore, the in-situ detection of IFN-γ is an important method in disease prevention, medical diagnosis, and biological engineering, etc.
Several optical IFN-γ biosensors have been researched in recent years, such as localized surface plasmon resonance (LSPR) [7], fluorescence immunoassays [8], and strand-displacement aptasensing strategy biosensors [9,10]. However, most of optical biosensors contain many optical devices and a microfluidic chip to construct a complicated optical system [11]. Besides, many optical biosensors, such as fluorescence, need complex digital signal processing (DSP) to interrogate the optical signal. Alternatively, label-free fiber optic biosensors based on immunoreactions have many merits, such as easy fabrication and integration. One hole in the hollow core fiber (HCF) can forms the optofluidic waveguide, which significantly simplifies the sensing configuration [12]. The detection target can be bound with the aptamer, which modulate the refractive index (RI) on the surface of the fiber. Many optical devices have been integrated into the optical fiber for the detection of the RI change, including fiber gratings in photonic crystal fibers (PCF) [13–15], the in-line interferometer in micromachined fibers [16,17], and the SPR in HCFs [18].
However, temperature cross-sensitivity is a long-standing challenge for each type of fiber optics biosensor. For fiber gratings, the temperature could change the RI or pitch of the fiber grating, causing a resonant wavelength shift. For fiber interferometers, temperature also changes the RI or effective length of the interferometer, which modulates the optical path difference of fiber interferometers significantly. For the SPR, the temperature could change the RI of the dielectric, producing a variation of the resonance condition. Therefore, the temperature cross-sensitivity must be eliminated in order to realize a fiber optics biosensor with high accuracy and high stability. Some researchers have been investigated some methods to compensate the temperature cross-sensitivity, such as the combination of different optical devices or the same temperature responses of two polarization modes [19,20]. However, some methods contain complex fabrication process or DSP, such as no core fiber combined with a fiber Bragg grating [19], some other methods may reduce the RI sensitivity of the sensor, such as phase shifted fiber Bragg grating on a side-hole fiber [20]. Hence the temperature compensation is still a challenge for the fiber optic biosensor.
In this paper, a dual-optofluidic waveguide antiresonant reflecting optical waveguide (ARROW) for the detection of IFN-γ concentration with temperature compensation has been proposed. Two cladding holes were filled with IFN-γ and sodium chloride (NaCl) solution, respectively, which form a dual-optofluidic waveguide ARROW. The IFN-γ concentration can be detected through the modulation of the lossy mode resonance condition for the ARROW induced by the immunoreaction between the aptamers and IFN-γ. Meanwhile, the wavelength shift of two resonance dips are the same to each other based on two ARROWs with the similar thermo-optic coefficient, and the temperature cross-sensitivity can be compensated effectively. Hence, the accuracy and stability of the fiber optic biosensor could be improved significantly, which has a great potential for the real-time detection of IFN-γ concentrations in the fields of health monitoring, cancer prevention, biological engineering, etc.
2. Fabrication method of the sensor
In the proposed sensor, a HCF was employed as the sensing fiber, as shown in Fig. 1(a). The HCF is consisted of an air octagon core, an air-ring cladding with eight holes, and a silica cladding. The geometric size is shown in Fig. 1(b). Firstly, one end of the HCF was spliced with a section of the single mode fiber (SMF) in order to block all the holes. Then the SMF was cut using femtosecond laser irradiation, and the length of the remaining SMF was ∼10 µm. After that, a channel was fabricated inside the SMF by using the femtosecond laser in order to connect the outside and one cladding hole in the HCF. After that, the NaCl solution was filled into a cladding hole with a length of 10 cm through the capillary force by immersing the remaining SMF into the NaCl solution for ∼48 h, as shown in Fig. 1(c). A short section of the HCF with the SMF was removed by using a fiber cleaver. Finally, two SMFs were then spliced with two ends of the NaCl-infiltrated HCF.
Fig. 1. (a) The cross section of the HCF. (b) Geometric size of the HCF. (c) The cross-section of the NaCl-infiltrated HCF. (d) The close-up of the microchannel. (e) The schematic diagram of the dual-optofluidic waveguide ARROW.
Then in the silica cladding of the HCF, two square microchannels were drilled in the silica cladding through the femtosecond laser micromachining. Noted that two microchannels should align to the other hole in the air-ring cladding, as illustrated in Fig. 1(d). The liquid sample could be injected into the hole of the HCF through one microchannel, as an inlet. Then the liquid sample could be also pumped out of the hole through the other microchannel, as an outlet, as shown in Fig. 1(e). As a result, an optofluidic waveguide was formed inside the HCF. In this experiment, we fabricated two biosensors with the same structure of optofluidic waveguide mentioned above. One was used for the measurement of the IFN-γ with single-stranded aptamer DNA, the other one was used for the measurement of the IFN-γ with signal-enhanced streptavidin- aptamer layer
Theoretically, the NaCl can be infiltrated into any hole in the air ring cladding, even two adjacent channels. However, due to the thin thickness of the fiber skeleton in the HCF (∼2.5µm), too much liquid concentrated in two closed holes may cause a deformation or even damage of the fiber skeleton. Hence in this experiment, two channels in the air ring cladding with the interval of two holes was selected as the optofluidic waveguide and referenced ARROW, as shown in Fig. 1(e).
3. Principle of the proposed sensor
Figure 2(a) shows the principle of the HCF, which can be described as an ARROW model [21]. Due to the high refractive index of the IFN-γ or NaCl infiltrated cladding in the HCF, the guided light can be reflected at two interfaces of the cladding, forming a Fabry-Perot resonator, as illustrated in Fig. 2(b). The guided light at the resonant wavelengths leaks out of the cladding of the HCF, resulting in the periodic and narrow resonance dips in the transmission spectrum [22]. Hence an ARROW is formed in the HCF. However, there are two ARROW in the HCF because of two different infiltration materials: the IFN-γ solution, and the NaCl solution. Figures 2(c) and 2(d) illustrated the numerical simulations of light distribution. At the wavelength of 1550.38nm, the resonance condition of the ARROW for the optofluidic waveguide is achieved, as shown in Fig. 2(c). On the other hand, at the wavelength of 1557.86 nm, the resonance condition of the ARROW for the NaCl infiltrated channel is achieved, as shown in Fig. 2(d) [23]. Hence, there are two resonance dips corresponding to two materials, which can be expressed as [24]
Fig. 2. (a) Diagram of the dual-optofluidic waveguide ARROW and (b) the Fabry-Perot resonator. Numerical simulations of light distribution at the wavelength of (c) IFN-γ infiltrated resonator and (d) NaCl infiltrated resonator. (e) Experimental setup of the biosensor. (f) The fluorescent image of the optofluidic waveguide channel filled with fluorescent dye labelled aptamer layer.
The immunoreaction between the aptamer and the IFN-γ could modulate the wavelength of the resonant dip for the IFN-γ infiltrated ARROW due to the changed RI inside the optofluidic waveguide. However, the wavelength of the resonant dip for the NaCl infiltrated resonator is fixed because of the independent channel. On the other side, wavelength shifts of two resonance dips corresponding to two ARROWs have the same response to the temperature fluctuation because two infiltration materials have the similar thermo-optic coefficient. Therefore, a dual- optofluidic waveguide ARROW biosensor is formed with the temperature compensation by interrogating the wavelength interval between two resonance dips of the dual- optofluidic waveguide ARROW.
Because the elimination of the temperature cross-sensitivity for the biosensor is used for the similar temperature responses of two infiltration materials in two optofluidic waveguide ARROWs, the thermo-optic coefficients of two infiltration materials must be same to each other. In order to eliminate the temperature cross-sensitivity of the biosensor, the temperature responses of the NaCl and running buffer were investigated. We measured the refractive index of the NaCl and running buffer (RB) by using a ATAGO refractometer, as shown in Fig. 3. The thermo-optic coefficient of the RB is about -3.04 × 10−4/°C. Then the temperature response of the NaCl at different concentrations was also measured. The thermo-optic coefficient of the NaCl with a concentration of 30% is also about -3.10 × 10−4/°C, which is close to that of the RB. Therefore, the NaCl with the concentration of 30% was chosen in order to eliminate the temperature cross-sensitivity. Noted that the RB is the best choice for the referenced ARROW because of the exactly matched thermo-optic coefficient. However, the same refractive index of two liquid in the optofluidic waveguide and referenced ARROW results the overlap between two resonance dips in the transmission spectrum. Although many materials immobilized on the surface of the optofluidic waveguide, such as Poly-L-lysine or Single-stranded aptamer DNA, could also make a wavelength difference between two resonance dips, the wavelength difference (∼1.32nm in section 4.4) is smaller than the full width at half maximum (FWHM) of the resonance dip of the optofluidic waveguide (∼1.44nm in section 4.4). Especially in the measurement of IFN-γ at low concentration, most part of two resonance dips would be overlapped to each other, which is very hard to distinguished in the experiments. Therefore, the NaCl with the concentration of 30% was chosen for both the similar thermo-optic coefficient and the enough wavelength difference between two resonance dips in the transmission spectrum.
4. Experiment and discussion
4.1 Experimental setup
Figure 2(e) shows the experimental setup. Light from an amplified spontaneous emission (ASE, 1525nm∼1565nm) was launched into the fiber and the transmission spectrum was recorded by an optical spectrum analyzer (OSA, AQ-6370D, Yokogawa Co., Ltd.) with a resolution of 0.02nm. A polarizer and a polarization controller (PC) were placed between the ASE and the ARROW. The HCF was slightly pre-stretched and fixed in a metal plate. The liquid sample is pumped into the optofluidic waveguide channel with two microchannels through the covered solidified polydimethylsiloxane (PDMS).
4.2 Materials
Recombinant human IFN-γ was purchased from Cyagen Inc. Dye labeled Single-stranded aptamer DNA (5’ -GGGAGGTTCGTGGTACTATTCGGGCGGTGT-3’) was synthesized by Shanghai Sangon Biotech. Tumour necrosis factor alpha (TNF-α), interleukin-2 (IL-2), interleukin 6 (IL-6) were purchased from ACRO Biosystems. Poly-L-lysine (pH 7.4, 0.1 (w/v)%) (PLL), sodium chloride (NaCl), running buffer (RB) (NaH2PO4 pH 7.4, 150 mM NaCl), streptavidin, toluene, phosphate buffered saline (PBS) were bought from Sigma-Aldrich Pty. Ltd. Femtosecond laser was purchased from Spectra Physics, Inc. Biotin-PEG-silane was got from Melopeg Inc.
4.3 RI and temperature response of the dual-optofluidic waveguide ARROW biosensor
Figure 4(a) shows the transmission spectrum of the ethanol and NaCl infiltrated dual-optofluidic waveguide ARROW biosensor. There are two resonance dips in the wavelength range from 1525 nm to 1565 nm. The resonance dips at the wavelength of 1545.58 nm and 1550.86 nm are in good agreement with the theoretical predictions of an ethanol-infiltrated Fabry-Perot resonator (1545.78 nm) and NaCl-infiltrated Fabry-Perot resonator (1550.38 nm). Therefore, two resonance dips are generated simultaneously with two materials-infiltrated Fabry-Perot resonators.
Ethanol solutions with different RI ranged from 1.3568 to 1.3622 RIU (corresponding to different concentrations) (ATAGO refractometer) were injected into the optofluidic waveguide channel. Figure 4(b) shows the wavelength shift of the dual-optofluidic waveguide ARROW with different RI at the temperature of 20°C. The resonance dip corresponding to the ethanol-infiltrated Fabry-Perot is shifted to a longer wavelength due to the increase of RI. However, the resonance dip corresponding to the NaCl -infiltrated Fabry-Perot remains constant because the RI of the NaCl is fixed. The wavelength of two resonance dips with different RI was plotted in Fig. 4(c). Besides, the wavelength interval between two resonance dips is also shown in Fig. 4(d). The sensitivity of -1413 nm/RIU can be achieved for the RI response with the resolution of 0.02nm of the OSA in this experiment. Hence, the optofluidic waveguide channel can be used as a biosensor to detect the immunoreaction between the aptamer and IFN-γ due to the modulation of the wavelength interval with the small RI change. The error bars in Fig. 4(d) are around 0.3-1.3 nm. Although the temperature influence has been already eliminated, the optofluidic waveguide is also suffered from other cross-sensitivity. For example, the air turbulence may lead a micro bending for the HCF, making a wavelength shift of the resonance dip. Besides, the flow of the liquid sample makes a slight deformation of the optofluidic waveguide, also causing a wavelength shift of the resonance dip.
Fig. 4. (a) Transmission spectrum of the dual-optofluidic waveguide ARROW biosensor. (b) Wavelength shifts with different RI. (c) Relationship between the wavelengths of resonance dips and RI. (d) Wavelength interval with different RI. (e) Wavelength shifts at different temperature from 20 to 80°C. (f) Wavelength of resonance dips and the wavelength interval.
Figure 4(e) illustrates transmission spectra of the biosensor at different temperatures from 20 to 80°C. Two resonance dips are shifted to longer wavelengths simultaneously at different temperature. However, the wavelength interval is kept constantly with the standard variation of 0.02 nm due to the same temperature response of two ARROWs, as plotted in Fig. 4(f). Hence, by interrogating the wavelength interval, the dual- optofluidic waveguide ARROW biosensor becomes insensitive to the temperature.
4.4 The measurement of the IFN-γ with single-stranded aptamer DNA
Aptamer is a kind of short sequence oligonucleotides with a specific three-dimensional structure [25]. Through the artificial synthesis, the aptamers could bind their targets with high affinity and specificity. Aptamer possesses many unique merits, including convenient commercial synthesis processes, good consistency, and fast response for real-time biomarkers detection [26]. The aptamer-based biosensors are defined as aptasensors, such as the IFN-γaptasensors. Generally, the aptamers are immobilized on the surface of the aptasensors. The immunoreaction between the aptamers and the IFN-γ make a change for one specific parameter, such as current, fluorescence emission, or refractive index [27]. The aptasensors could detect the IFN-γ through the parameter change. In the proposed fiber optic biosensor, the RI change induced by the binding of aptamers and the IFN-γ could modulate the resonance condition of the ARROW.
The first step for the measurement of IFN-γ is the immobilization of the single-stranded aptamer DNA inside the optofluidic waveguide channel. The optofluidic waveguide channel was injected with piranha solution and washed with PBS to clean the optofluidic waveguide channel. Then a PLL solution was pumped into the optofluidic waveguide channel for 20 min to form a PLL thin layer with an amino group because the silica surface of the optofluidic waveguide channel is negative charged. The optofluidic waveguide channel was rinsed again with a RB in order to clean the unbound PLL. PLL on the surface of the optofluidic waveguide with extreme positive charge could bind the aptamer inside the optofluidic waveguide channel due to the negative charge of the single-stranded aptamer DNA [28]. A 20 ml of 100 nM single-stranded aptamer in the buffer was injected into the optofluidic waveguide channel. After standing for one hour, the single-stranded aptamer was bound to the PLL and formed a uniform thin layer inside the optofluidic waveguide channel. The unbound aptamer was also washed with the rinsing of RB. The detail of each modification is shown in Fig. 5(a).
Fig. 5. (a) The detail of each modification process. (b) Transmission spectra of biosensor before (dash line) and after washing (solid line) for each modification. (c) Wavelength of resonance dips and the wavelength interval. (d) Wavelength interval change with different IFN-γ concentration.
Figure 5(b) shows the wavelength shift of two resonance dips before (dash line) and after (solid line) washing for each modification. Compared with the ethanol-infiltrated Fabry-Perot resonator shown in Fig. 4(a), the resonance dip corresponding to the RB-infiltrated ARROW is shifted to 1555.38 nm. Figure 5(c) illustrate the wavelength shifts of two resonance dips and the wavelength interval. The resonance dip experiences a red shift (from 1555.38 to 1557.04 nm) as a consequence of the adsorption of an IFN-γ layer on the surface of the ARROW. After immobilization of the single-stranded aptamer DNA, an IFN-γ solution of 2.0 ng/ml was also pumped into the optofluidic waveguide channel. Figure 2(f) shows the fluorescent image of the dual-optofluidic waveguide ARROW biosensor filled with fluorescent dye labelled IFN-γ layer (fluorescence excitation at 494 nm and emission at 526 nm), which prove that a uniform aptamer layer was formed inside the optofluidic waveguide. The wavelength of the resonance dip shifts from 1557.04 nm to 1557.96 nm after the washing of the un-associated IFN-γ molecules by using the RB due to the immunoreaction between the aptamer and IFN-γ. However, the other resonance dip corresponding to the NaCl-infiltrated Fabry-Perot has been fixed during the entire modification and associated events. Thus, the wavelength interval is changed as 0.92 nm, indicating that the IFN-γ solution can be detected by interrogating the wavelength interval.
Figure 5(d) illustrates the wavelength interval change during the entire measurement with different concentrations of IFN-γ solution ranging from 0.5 ng/ml to 2000mg/ml. It was found that the biosensor can be regenerated for 12 detection/regeneration cycles. The result was fitted by OriginPro 8 using the Dose Response Analysis method. The threshold (dose) concentration and the maximal effect of the curve were measured as 0.9 ng/ml and 200 ng/ml, respectively, indicating that the limit of detection (LOD) is roughly 0.9 ng/ml and the maximum binding capacity is around 200 ng/ml. With the concentration range from 2 ng/ml to 200 ng/ml, a response sensitivity of 0.039 nm/(ng/ml) can be achieved. It should be noted that the RI response of the sensor is also dependence of the HCF length because the amount of the aptamer on the surface of the optofluidic waveguide is rely on the length of the HCF. Hence a proper HCF length should be adjusted in the experiment.
4.5 The measurement of the IFN-γ with signal-enhanced streptavidin-aptamer layer
The optofluidic waveguide channel in the other biosensor was injected into piranha solution to repeat the same cleaning process. Besides the cleaning of the optofluidic waveguide, the piranha solution can be also used for the hydroxylation step. On the surface of the optofluidic waveguide (silica), amount of silicon-oxygen bond is existed. The piranha solution (acid) could react with the silicon-oxygen bond to generate the hydroxyl group. The Biotin-PEG-silane solution was then pumped into the optofluidic waveguide channel for six hours in order to immobilize the Biotin-PEG-silane linkers on the surface of the optofluidic waveguide channel with the hydroxyl group. The optofluidic waveguide channel was washed with toluene, acetone, and deionized water in turn for two hours to clean the non-covalently attached Biotin-PEG-silane linkers. Subsequently, the streptavidin solution and single-stranded aptamer DNA were injected into the optofluidic waveguide channel to form a uniform aptamer layer, followed by a 20 min washing with DI water and RB.
The streptavidin was bound with a Biotin-PEG-silane linker immobilized on the surface of the optofluidic waveguide through hydroxyl group, as shown in Fig. 6(a). The streptavidin was served as the bridge which easily linked with the aptamer. In each streptavidin, four biotin-binding points exist simultaneously. Hence, three strands of single-stranded aptamer DNA can be linked with each Biotin-PEG-silane linker, making a 3-fold signal amplification. Figure 6(b) illustrates the wavelength interval change of the dual-optofluidic waveguide ARROW functioned with streptavidin with the IFN-γ concentrations range from 0.1ng/ml to 8000 ng/ml. Obviously, the wavelength interval change is significantly larger than that without the streptavidin, indicating the signal enhancement effect of the streptavidin. The LOD is 0.5ng/ml and the maximum binding capacity is around 1000ng/ml, respectively. The response sensitivity of 0.088nm/(ng/ml) can be achieved for the signal-enhanced biosensor, which is 2.25 times larger than that without the streptavidin. The experimental result also confirms the signal amplification of the streptavidin.
Fig. 6. (a) The detail of each modification process. (b) Wavelength interval change with different IFN-γ concentration for two methods. (c) Time response of the biosensor. (d) Wavelength interval change with IFN-γ and different proteins.
Figure 6(c) shows an enlarged view of the wavelength interval change with the IFN-γ concentration of 5ng/ml. After the injection of IFN-γ solution, the wavelength interval was changed by 5.56 nm in early 66 s. After that the optofluidic waveguide channel was washed to remove the unbound IFN-γ. Hence at 160s the wavelength interval change is decreased to 4.58 nm due to the washing of the unbound IFN-γ. Then the wavelength interval change is dropped for 92 s when the optofluidic waveguide channel was washed with 1 mM HCl, ultrapure water, and RB, indicating that the time response of the majority association and dissociation procedure is occurred in 66 s and 92 s.
The specificity of the dual-optofluidic waveguide ARROW was also researched. Five different proteins were pumped into the optofluidic waveguide channel in sequence with the concentration of 0.3 pM, as illustrated in Fig. 6(d). The cross-sensitivity wavelength shift of other proteins is caused by the blank space on the surface of the optofluidic waveguide. Although the BSA can fill the blank space, there are still some small blank space, and IFN-γ can be dropped into the blank space, resulting a cross-sensitivity wavelength shift of the biosensor. The maximum cross-sensitivity of 27.6% is achieved with the IgG among five proteins, showing a good specificity of the aptamer-based IFN-γ biosensor.
In order to research the reusability of the sensor, IFN-γ with concentrations of 5ng/ml were injected into the optofluidic waveguide repeatedly. In each measurement, the solution was kept for 20 s in order to allow all the IFN-γ in the solution to bind completely to the immobilized aptamer, then followed by a 60 s washing in ultrapure water. After each measurement, the optofluidic waveguide channel was injected into the HCl, ultrapure water, and RB to remove the IFN-γ. Figure 7 shows the response of the sensor with the concentration of 5ng/ml. It can be seen that in the initial 12 measurements, the wavelength interval change remained at almost 4.58 nm. However, for the subsequent 6 measurements, the wavelength interval change began to decrease sequentially. After the 19th measurement, the wavelength interval change was almost fixed without any change, indicating the degradation of the aptamer induced by the HCL. Hence the reusability of the proposed sensor is 18 times.
4.6 The temperature compensation of the dual-optofluidic waveguide ARROW biosensor
The temperature fluctuation is always a serious cross-talk for most optical biosensors, especially for the fiber optic sensors. The proposed dual-optofluidic waveguide ARROW biosensor was fixed in the environmental chamber to investigate the temperature response of the biosensor. In the optofluidic waveguide, IFN-γ with a concentration of 2 ng/ml was immobilized on the surface. The temperature of the environmental chamber was adjusted from 20°C to 80°C with an interval of 10°C. Figures 8(a) and 8(b) show the wavelength of two resonance dips at different temperatures. Two resonance dips for the IFN-γ and NaCl-infiltrated resonators are shifted to longer wavelength simultaneously due to the RI change at the temperature range from 20°C to 70°C. However, the sensitivity response of two resonance dips is the same for each of them due to the same thermo-optic coefficients of two liquids and the same geometrical size of the two optofluidic waveguide channel. Hence, the wavelength interval between two resonance dips is almost fixed with only a slight variation, as shown in Fig. 8(c). It should be noted that the wavelength interval change decreases when the temperature exceeds 70 °C. At high temperature, the physical or chemical characteristic of protein could be changed. Hence the aptamer and IFN-γ become denatured when the temperature exceeds 70 °C. As a result, the aptamer may fall out of the surface of the optofluidic waveguide. Even the aptamer is still immobilized on the surface of the optofluidic waveguide, the IFN-γ can’t bind with the aptamer due to the change of physical or chemical characteristic for the aptamer or IFN-γ. Hence the wavelength interval change decreases when the temperature exceeds 70 °C. Figure 8(d) shows the temperate response of the wavelength interval change with different IFN-γ concentrations. In the temperature range from 20 to 70°C, the maximum standard variation is 0.06 nm with the IFN-γ concentration of 1000 ng/ml. Thus, the dual-optofluidic waveguide ARROW could eliminate the temperature cross-talk very effectively.
Fig. 8. Wavelength shifts of the resonance dip at the wavelength of (a) IFN-γ infiltrated resonator and (b) NaCl infiltrated resonator. (c) Wavelength of two resonance dips and the wavelength interval. (d) Wavelength interval change with different IFN-γ concentration.
Table 1 provides a summary of the performance of the fiber IFN-γbiosensors with similar sensing configuration. In general, the LOD of the proposed optical fiber sensor is lower than most of other aptasensors, but higher than that based on the of fluorescence aptasensor of 0.25 ng/mL. Besides, the operation range of the proposed optical fiber sensors are larger than that of other aptasensors. More importantly, the proposed ARROW IFN-γbiosensor could realize the temperature compensation, which is hard to be achieved for other aptasensors.
Table 1. Performance of fiber optics sensors with similar sensing configuration.
5. Conclusion
In conclusion, a dual-channel ARROW biosensor for the detection of IFN-γ concentration with temperature compensation was proposed. Two Fabry-Perot resonators, infiltrated with IFN-γ and NaCl, were formed in the HCF, which generated two resonance dips in the transmission spectrum based on the ARROW model. The IFN-γ concentration can be measured by interrogating the wavelength interval between two resonance dips. Meanwhile, the temperature fluctuation can be eliminated simultaneously due to the same thermo-optic coefficients of two infiltration liquids. Two kinds of IFN-γ detection have been carried out, one with the single-stranded aptamer DNA, and the other with the signal-enhanced streptavidin-aptamer layer. The experimental results show that the streptavidin can amplify the biosensor response significantly, and the LOD of the IFN-γ concentration of 0.5 ng/ml is achieved. More importantly, the influence of the temperature could be compensated through the referenced resonance dip, and so the accuracy of the IFN-γ detection can be improved significantly. The proposed fiber biosensor can be used for the real-time measurement of IFN-γ concentration in the fields of health monitoring, cancer prevention, and biological engineering, etc.
Funding
National Key Research and Development Program of China (2019YFA0706304); National Natural Science Foundation of China (61601436, 61675033, 61727817, 61835002).
Disclosures
The authors declare no conflicts of interest.
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