Optofluidic immunosensor based on resonant wavelength shift of a

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Optofluidic immunosensor based on resonant wavelength shift of a hollow core fiber for ultratrace detection of carcinogenic Benzo[a]pyrene Ran Gao, Dan-feng Lu, Mengying Zhang, and Zhi-Mei Qi ACS Photonics, Just Accepted Manuscript • DOI: 10.1021/acsphotonics.7b01009 • Publication Date (Web): 22 Feb 2018 Downloaded from http://pubs.acs.org on February 25, 2018

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ACS Photonics

Optofluidic immunosensor based on resonant wavelength shift of a hollow core fiber for ultratrace detection of carcinogenic Benzo[a]pyrene Ran Gao, Dan-Feng Lu, Meng-Ying Zhang, and Zhi-mei Qi State Key Laboratory of Transducer Technology, Institute of Electronics, Chinese Academy of Sciences, Beijing 100190, China

ABSTRACT Polycyclic aromatic hydrocarbons (PAHs) are always regarded as contaminants originating from the incomplete combustion of organic compounds. Benzo[a]pyrene (B[a]P) can be adopted as a marker for the overall PAH mixture due to its ultra-toxic property, which can cause cancer of the lungs, skin, and prostate. It is necessary to monitor the B[a]P contamination levels in a water environment. In this paper, an in-line fiber optofluidic sensor for the detection of B[a]P in water by using an anti-resonant reflecting optical waveguide is proposed and experimentally demonstrated. One air hole in the cladding of the hollow-core fiber was fabricated as an in-line fiber optofluidic combined with two inlet and outlet micro-channels fabricated by using femtosecond laser micromachining. The B[a]P molecule can be detected through the anti-resonant reflecting optical waveguide due to the immunoreaction between the antibody and the target B[a]P. The experimental results show that a sensitivity of up to 23 pm/pM is achieved. The proposed fiber sensor appears to have potential applications in research on environmental contamination, food safety, human health, etc.

Keywords:

Anti-resonant

reflecting

optofluidic;

Hollow

micromachining fabrication; B[a]P detection.

ACS Paragon Plus Environment

core

fiber;

Femtosecond

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TABLE OF CONTENT GRAPHIC


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Polycyclic aromatic hydrocarbons (PAHs), which are formed from at least two condensed aromatic rings composed of carbon and hydrogen atoms, can easily cross cell membranes and tend to bioaccumulatein lipid tissues [1]. In general, PAHs are identified as carcinogens that cause cancer of the lungs, skin, prostate, and act as endocrine-disrupting compounds [2]. PAHs are generally produced from incomplete combustion of organic material, or improper disposal of fuels and oils [3]. As one of the most toxic PAHs, benzo[a]pyrene (B[a]P) can be adopted as a marker for the overall PAH mixture [4]. Due to the extreme toxic effect of B[a]P, many organizations have ruled on a rigid upper limit, such as 200 ng/mL for the United States Environmental Protection Agency (US EPA) [5], 10ng/L in the European Drinking Water Directive (98/83/EC) [6], and 10ng/L for the National Health and Family Planning Commission of the People’s Republic of China [7]. Therefore, monitoring of B[a]P plays a critical role in many fields, such as environmental contamination, food safety, human health, etc. Immunoassays have been widely applied in many biosensors based on the immunoreactions between the antibody and the target in the antigen. For PAH, various immunoanalytical methods have been developed, such as electrochemical immunosensors [8, 9], piezoelectric transduction [10], enzyme-linked immunosorbent assays (ELISA) [11, 12], etc. Compared with several conventional immunosensors, optical biosensors have shown potential for B[a]P quantification detection due to their significant merits, such as label-free detection, high sensitivity, and immunity to electromagnetic fields. Therefore, a series of optical biosensors for the detection of B[a]P have been developed, including polarization modulation reflection absorption infrared spectroscopy (PMRAIRS)[13], quartz crystal microbalance (QCM) biosensors [14], surface IR immunosensors [15], fluorescence immunoassays [16], and surface-plasmon resonance (SPR) biosensors [17]. In general, conventional optical biosensors always need microfluidic on-chip waveguides combined with other optical components, including a light source, lens, and detectors [18]. However, in real applications, analytical biosensors are always required with small size and easy operation. The hollow core fiber in-line optofluidic is a good alternative due to the hollow holes inside the fiber. Various optical fiber components could be fabricated into the fiber, such


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as fiber gratings (i. g. fiber Bragg gratings (FBGs) or long period fiber gratings [20].) or in-line interferometers (Fabry–Pérot interferometer [21] or Sagnac interferometer [22].), The small refractive index change in the in-line optofluidic could modulate the resonant condition of the fiber gratings or optical path difference (OPD) of the fiber interferometer. The fiber in-line optofluidic biosensors possess high sensitivity, low cost, and immune to electromagnetic interference. The anti-resonant reflecting opticalwaveguide (ARROW) model has attracted much attention since its first demonstration [23]. Due to the low refractive index of the hollow core, some part of the guide light is reflected at the interface between the hollow core and the cladding. The other part of the guide light is transmitted through the cladding, and reflected at the interface between the cladding and outside air.Hence the guide light is reflected at the inside and outside interfaces of the cladding, which is considered as a Fabry–Pérot etalon [24]. The ARROW of the capillary waveguide is sensitive to the surrounding environments, and various sensing applications have been proposed, includingtemperature [25], magnetic field [26], displacement [27],water-level sensing [25], and so on. In this paper, we proposed an in-line fiber optofluidic biosensor for the detection of B[a]P based on the guidance mechanism of the ARROW. The hollow hole in the cladding of the hollow-core fiber (HCF) forms an in-line optofluidic channel. The B[a]P solution can be flowed into the channel through the inlet and outlet port fabricated by using the femtosecond laser micromachining. The immunoreactions between the antibody and the B[a]P leads a significant change of the refractive index inside the in-line optofluidic channel, and the resonant condition of the fiber cladding is modulated, as a double-layered Fabry–Pérot resonator. The concentration of the B[a]P molecular can be measured by interrogating the wavelength shift of the lossy dip. The proposed in-line fiber optofluidic B[a]P biosensor appears to have potential applications in research into environmental contamination, food safety, human health, etc.

MATERIALS AND METHODS MATERIALS


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Sulfuric acid (H2SO4), hydrogen peroxide aqueous solution (H2O2), and hydrogen chloride (HCL)

were

purchased

from

ShangHaiHushi

Laboratorial

Equipment

Co.,

Ltd.

8-aminopyrene-1,3, 6-trisulfonic acid, trisodium salt (ATPS), glutaricdialdehyde, and Bovine Serum Albumin (BSA) solution were purchased from Sigma-Aldrich. Phosphoric acid buffer solution (PBS) was obtained from HyClone Laboratories. B[a]P mouse monoclonal antibody and goat anti-mouse IgG conjugated with fluorescein isothiocyanate (FITC) were bought from Santa Cruz Biotechnology, Inc., while B[a]P and other PAHs standard solution were purchased from the National Institute of Metrology (China). All solutions were prepared with ultrapurewater (18.4MΏ cm, Millipore Corp.).

Fabrication of the proposed in-line fiber optofluidic sensor In the proposed sensor, the air core of the HCF is a hollow octagon with a side length of 18µm. The cladding is consisted of eight hollow holes with an inner diameter of 35 µm, which form an air-ring. The diameter of the core, air-ring, and outer cladding is 36, 90, and 190µm, respectively, as shown in Figure 1(a). The schematic diagram of the proposed sensor is shown in Figure 1(b). A short section of the HCF with length of ~14 cm was manually fusion spliced between two standard SMFs with the gap of 10 µm, arc power of 18 mA, and arc duration of 140 ms to avoid the collapse of the air holes in the HCF. Then two micro-channels were drilled through the surface of the HCF. A femtosecond laser (5W, Spectra-Physics) with a central wavelength of 800 nm, pulse width of 35fs and repetition rate of 1kHz is employed for the sensor fabrication. The laser pulse energy is attenuated through a half-wave plate and a polarizer. Then a neutral density filter is applied to reduce the pulse energy before the objective lens. The attenuated femtosecond laser beam is focused by an objective lens (20×, NA=0.6, Olympus) and the focal spot diameter is about 3µm. Two square micro-channels with a side length of 24µm were fabricated on the surface of the HCF with a separation of 10 cm by controlling the movement of the stage, as shown in Figure 1(c). In this way, one hollow hole in the cladding and two micro-channels can be served as an in-line microfluidic channel, inlet, and outlet, respectively.


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Figure 1(a) Cross-section of the HCF. (b) The schematic diagram of the proposed sensor. (c) Square micro-channel for the inlet and outlet.

The experimental setup of the proposed in-line fiber optofluidic sensor The experimental setup is shown in Figure 2(a). An amplified spontaneous emission, of which the wavelength is 1525–1565 nm and the output power is 20 mW, was used as a broadband source to illuminate the fiber. The transmission spectrum of the sensor was interrogated by using an optical spectrum analyzer (OSA) (AQ6730B, Yokogawa Co., Ltd.). In order to protect the sensor, a metal plate was designed, as shown in the inset of Figure 2(a). In the metal plate, a V-groove channel was fabricated on the top surfaces of two holders to fix the fiber. The HCF was slightly pre-stretched and covered with two square solidified polydimethylsiloxane (PDMS) chips with a 1.5 mm diameter hole, which could seal the fiber to prevent liquid leakage. The liquid sample could pass into the PDMS chips through polytetrafluoroethylene (PTFE) tubes by using syringe pump.


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Figure 2 (a) The experimental setup of the proposed sensor. Inset shows the metal plate. (b) The optical micrograph of the fiber. Fluorescent image of the anti-mouse IgG -FITC filled fiber (c) all hollow holes, (d) only single hollow hole.

Immobilization of B[a]P antibody on in-line optofluidic First, the home-made piranha solution (H2O2:H2SO4 =in a 1:3 ratio by volume) was injected into one hollow hole in the cladding to clean off all organic matter through the syringe pump. After incubation for 2h, the piranha solution was pumped out of the optofluidic, and the ultrapure water was sequentially injected 3 times to wash until the pH of the cleaning water reached neutral. Abundant hydroxyl groups were formed on the surface of the optofluidic for the silylation. After that, 5% HCL was added into 5% APTS solution in order to adjust the pH of the APTS solution to 4. The APTS solution was filled into the optofluidic at 75° for 1 h. After rinsing with ultrapure water, the in-line optofluidic was dried in an oven at 60 °C for1 h. In this way, the surface of the optofluidic was silanised with the presence of amine groups. Finally, in order to bind the antibody to the surface of the optofluidic, the glutaricdialdehyde was prepared because the two aldehyde groups can be bound with the mercapto on the surface of the optofluidic and the amino in the antibody, as a cross-linking reagent. Hence the glutaricdialdehyde solution with a concentration of 2.5% was dissolved in the PBS buffer, which was deposited in the optofluidic for 1 h, and followed by rinsing in PBS 2 times and in ultrapure water 3 times. A 0.5mg/ml B[a]P mouse monoclonal antibody was flowed through the optofluidic for 36 h, followed by the ultrapure water washing process in order to remove the unbounded antibody. Although the B[a]P molecular is immobile on the surface of the in-line optofluidic through the immunoreaction with the antibody, it also can be bound at the antibody-unbounded space on the surface, which is hard to be removed by using the water. Thus a 2mg/ml BSA solution was further introduced as block protein to


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prevent non-specific binding of the B[a]P molecular on the surface of the optofluidic, followed by the ultrapure water washing process repeated 3 times. In order to confirm the immobilization of B[a]P antibody on the surface of the in-line optofluidic, the proposed sensor filled with the goat anti-mouse IgG conjugated with fluorescein isothiocyanate (FITC) was fabricated. On the other side, the other fiber filled with the goat anti-mouse IgG –FITC in all hollow holes by using capillary force was also fabricated. Both fibers infiltrated with one hole and all holes were washed by using ultrapure water after the injection of the goat anti-mouse IgG conjugated with FITC. Figure 2 (b)-(d) show the optical micrograph of the fiber, the fluorescent image of the anti-mouse IgG -FITC filled fiber in all hollow holes, and in only single hollow hole. Obviously, the fluorescence microscopy confirms that the liquid sample could pass through only one hollow hole in the fiber through two square micro-channels, which forms an in-line fiber optofluidic. More importantly, Figure 2 (d) shows that a uniform B[a]P antibody layer was distributed throughout the in-line optofluidic, which confirms the success of the immobilization of B[a]P antibody.

The principle of the in-line optofluidic Many previous works have already discussed the guidance mechanism of the ARROW. The capillary’s cladding region can be described as a Fabry–Pérot resonator, which allows the core modes to oscillate and radiate through the cladding. Due to the low refractive index of the hollow core, some part of the guide light is reflected at the interface between the hollow core and the cladding. Another part of the guide light is transmitted through the cladding, and reflected at the interface between the cladding and outside air. Hence the guide light is reflected at the inside and outside interfaces of the cladding, which is considered as a Fabry–Pérot etalon. It should be noted that both the radiation of the fundamental mode due to the low refractive index of the core and the Fabry–Pérot resonator of the cladding form the guidance mechanism of the ARROW, rather than the different explanations of the liquid-infiltrated HCPCF [23].

Then, the mode field distribution of the fiber was simulated by using commercial Comsol Multiphysics software. The refractive index of the liquid sample and silica were given by 1.33 and 1.45, respectively. Due to the mismatch of the core mode and cladding modes in the HCF, the fundamental core mode is confined in the air core, as shown in Figure 3 (a). However, when the


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liquid sample was flowed into the in-line optofluidic, the natural confine of the core mode in the fiber was broken. The refractive index of the core (air ~1.000) is less than that of the cladding. Thus, the fundamental core mode radiates and oscillates through the cladding of the liquid sample and silica, as shown in Figures 3 (b) and (c). The in-line optofluidic and the silica cladding are combined to form a double-layered Fabry–Pérot etalon [28, 29]. In the anti-resonant condition, the guided light is reflected back by the double-layered Fabry–Pérot etalon and propagated in the air core of the fiber due to the mismatching between the wavelengths and the resonant condition of the resonator. In the resonant condition, the guided light leaks out of the cladding of the fiber due to the matching between the wavelength and the resonant condition. Therefore, the principle of the in-line optofluidic can be described as the guidance mechanism of the ARROW. Thus, there is a series of lossy dips corresponding to the resonant condition in the transmission spectrum [28, 29]. Figure 3(d) and (e) show the mode field distribution at the wavelength of 1523.08 nm and 1534.16 nm when the in-line optofluidic is filled with liquid sample, which corresponds to the anti-resonant and resonant conditions. At the wavelength of 1523.08 nm, the guide light is confined in the air core in the fiber (shown in Figure 3 (d)), indicating the anti-resonant wavelength for the Fabry–Pérot resonator. In contrast, at the wavelength of 1534.16 nm, the mode field distribution leaks out of the air core and radiates into the liquid sample and silica cladding (shown in Figure 3 (e)). Note that in Fig. 3 (a) and (d), the mode field distributions are simulated as the fundamental mode of the waveguide. Without the liquid sample, the fundamental mode field is concentrated on the centre of the air core, and the boundary between the air core and cladding is weak. However, when the fibre is infiltrated with the liquid sample, the transmission window of the ARROW is modulated by the presence of the liquid in the channel, and the fundamental mode field is distributed in the air core almost uniformly.


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Figure 3 (a) Numerical simulation of the in-line optofluidic without the liquid sample. (b) Schematic diagram of the in-line optofluidic. (c) Guiding mechanism of the in-line optofluidic. Mode field distribution of the in-line optofluidic filled with liquid (d) at the anti-resonant wavelength (1523.08 nm), (e) at the resonant wavelength (1534.16 nm).

The wavelength of the lossy dips at the resonant condition, λr , can be expressed as [25]

λr =

2 2 2 2(d OP nLI2 − nair + d cl nsilica − nair )

m

. (1)

where m is the resonance order, dOP and d cl are the diameter of the hollow hole and the thickness of the outer silica, nair , nLI and nsilica are the refractive index of the air, liquid sample and the silica, respectively. The lossy dips at the resonate condition were simulated as 1523.08, 1534.16, and 1548.49, 1563.10 nm by calculating Eq. (1). Once the antibody is immobilized on the surface of the hollow hole, the immunoreactions between the antibody and the B[a]P will lead to a change in the refractive index in the in-line optofluidic, and the characteristic of the leaky modes is changed due to the interaction between the leaky modes and the B[a]P. The resonant condition of the Fabry–Pérot resonator is changed, and the lossy dips corresponding to the resonant condition are shifted. By interrogating the wavelength shift of the dip wavelength, the concentration of the B[a]P can be detected.

RESULTS AND DISCUSSION


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Measurement of immobilization of B[a]P antibody and association of B[a]P in the in-line optofluidic Since the presented in-line optofluidic fiber sensor is based on refractometry, sensitivity to refractive index changes should be measured and presented. A series of ethanol solutions with different concentrations were prepared as samples ranging from 1.3465 to 1.3632 RIU. The RI sensitivity of the in-line optofluidic is 1305.7 nm/RIU (see Supporting Information). The transmission spectra of the in-line optofluidic washed with ultrapure water after cleaning with piranha solution were investigated first. The transmission spectra of the in-line optofluidic with and without the liquid sample are shown in Fig. 4 (a). Obviously, the amplitude of the liquid -infiltrated in-line optofluidic (-12.2 dB) is reduced significantly compared with that without the liquid (-8.3 dB) due to the strong light absorption of the liquid. Meanwhile, four narrow lossy dips are shown in the transmission spectrum of the liquid- infiltrated in-line optofluidic, indicating the leaky effect. The wavelengths of the lossy dips are 1527.64, 1537.34, 1548.14 and 1558.86 nm, respectively, which are in good agreement with the theoretical predictions (1523.08, 1534.16, and 1548.49, 1563.10 nm).


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Figure 4 (a) Transmission spectrum of the in-line optofluidic with the liquid sample.(b) The transmission spectra of the sensor with the glutaricdialdehyde, antibody, BSA, and B[a]P. (c) The wavelength shift of the transmission spectra at various surface events after ultrapure water rinsing.

The experiments were carried out in situ and through the immobilization of B[a]P antibody on in-line optofluidics. Figure 4(b) shows the wavelength shift of one lossy dip in the transmission spectra of the modification and associated events after rinsing. The significant shift is induced by a change in the refractive index at each surface event before and after rinsing. The modification of layers on the surface of the in-line optofluidic results in the increasing of the refractive index of the liquid sample. Hence the initial resonance peak at 1537.34 nm undergoes a red shift during the experiments. After immobilization of the B[a]P antibody and BSA, a B[a]P solution of 100 pM was pass through the in-line optofluidic, followed by the repeated washing process with the ultrapure water in order to remove the unassociated B[a]P molecular. Figure 4 (c) shows the resonance wavelength shift of the lossy dip, and the insets depict the specific binding events. After injection with the 100 pM B[a]P and washing with ultrapure water, the lossy dip shifts from 1541.62 nm to 1545.74 nm, indicating the immunoreaction between the antibody and the B[a]P. It should be noted that each association follows a repeated process of wavelength shift after rinsing.


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This repeated process can be attributed to the rinsing of the unbound layer. Hence the rinsing of the in-line optofluidic is an accurate approach to eliminate the interference with unbound species.

The time and temperature response of the in-line optofluidic The time response of the sensor has been analyzed. A B[a]P solution with a concentration of 100 pM was injected into the in-line optofluidics and the lossy dip at the wavelength of 1537.34 nm was recorded for 400 s with the interval of 2 s. The measured result is shown in Figure 5 (a). From 0 to 56 s (before the first red dash line in Fig. 5(a)), the in-line optofluidic was washed with ultrapure water, hence the baseline was chosen as 1537.4 nm. Then the B[a]P solution was passed into the in-line optofluidic from 56 to 110 s, and the lossy dip shifted 4.78 nm. After that, the B[a]P solution was incubated in the in-line optofluidic for 60 s (from 110 to 170s). During this incubation time, the lossy dip continued to shift to 4.94 nm from 110 to 122s, indicating that the majority association procedure occurred in early 66 s (122s-56s) (between the first and second red dash lines in Fig. 5(a)). The wavelength shift of the lossy dip remains at 4.94 nm until 170 s. Then the wavelength shift begins to decrease to 4.12 nm at 182 s (between the third and fourth red dash lines in Fig. 5(a)) with the washing of the unbound B[a]P. After that the pepsin solution (50 ppm in glycine HCl buffer; pH 2.0) was injected into the in-line optofluidic in order to dissociated the B[a]P followed by the incubation of 30s, and the wavelength shift increase from 4.12 nm to 4.43 nm at 240 s (between the fifth and sixth red dash lines in Fig. 5(a)). The reason for the increasing of the wavelength shift is attributed to the high refractive index of the pepsin solution (50 ppm in glycine HCl buffer; pH 2.0; refractive index ~1.35) compared with that of ultrapure water. Finally, the in-line optofluidic was washed by using ultrapure water for 80 s (after the sixth red dash lines in Fig. 5(a)). The wavelength shift decreases to baseline again, and the dissociation time is about 54s (between the sixth and seventh red dash lines in Fig. 5(a)).


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Figure 5(a) The time response of the in-line optofluidic. (b) The temperature response of the in-line optofluidic. (c) Comparison of response of in-line optofluidic to 100 pM B[a]P and other PAHs. (c) Wavelength shift for the water sample.

The in-line optofluidic was fixed in a chamber, and the temperature dependence was investigated by adjusting the temperature range of the chamber from 20°C to 80°C. The photodegradation of the B[a]P makes it very volatile under sunlight. Especially at high temperature, the diffusion and collision of B[a]P molecules and oxygen are accelerated significantly. Hence the high temperature is beneficial for the generation and decomposition of the electricity transfer compound (CTC), which may aggravate the degradation of the B[a]P molecules [30]. Although the intensity of room light is weaker than that of sunlight, the photodegradation of the B[a]P molecules still has an impact on the accuracy of the in-line optofluidic. Therefore, the temperature experiment was carried

out in a dark room in order to avoid degradation of the B[a]P molecule at high temperature. Figure 5(b) shows the wavelength shift of the in-line optofluidic at different temperature. The temperature sensitivity of the wavelength shift is 2.8 pm/°C. Generally, in a laboratory environment the wavelength change induced by the temperature fluctuant (less than 3°C) is less than 8.4 pm. Compared with the concentration sensitivity of 23pm/pM for the B[a]P solution, the


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temperature cross-sensitivity can be neglected. Most of optical biosensors are very sensitive with the temperature, especially the fiber optic biosensors. However, the temperature cross-sensitivity of the proposed sensor is much less than that based on the method of the SPR or piezoelectric transduction [8, 17].

Selectivity of the in-line optofluidic for the detection of B[a]P To further verify the selectivity of the in-line optofluidic for the detection of B[a]P, 7 different PAHs with the same concentration of 100 pM were injected into the in-line optofluidic in sequence. These PAHs were chosen to investigate the selectivity of the sensor based on their similar molecular weight, which may interfere with the selective association of the B[a]P antibody. After each measurement, the in-line optofluidic was washed with ultrapure water for 10 s until the wavelength of the lossy dip shifted back to the baseline. The measured results are shown in Figure 5(c). The cross-reactivity rate is expressed as R = λPAH / λB[ a ]P × 100%. . Benzo[b]fluoranthene and benzo[a]anthracene led to a dramatic wavelength shift for the lossy dip of the sensor. However, the other four PAHs only induced a small wavelength change compared with that of the B[a]P (cross-reactivity rate of 3.93% for benzo[ghi]perylene, 2.91% for phenanthrene, 5.33% for benzo[k]fluoranthene, and 0.52% for fluorene), indicating the good selectivity of the in-line optofluidic for the detection of B[a]P. The small cross-reactivity may be explained by the similar number of condensed aromatic rings. Actually, the cross-reactivity is a common problem for the biosensors for the detection of the B[a]P [31]. The biosensors are hard to distinguish the chemical component of the detected substance. However, the immunoreactions between the antibody and the target B[a]P is an alternative method for the selective detection of B[a]P. In the proposed biosensor, the maximum cross-reactivity is only 24.1% due to the same

molecular weight between the B[a]P and B[b]F (252g/mol). Hence the immunosensor could improve the selectivity of the B[a]P detection significantly compared with other biosensors.

Immunoreaction between the antibody and the B[a]P at different concentrations A series of B[a]P solutions with different concentrations of 30 pM, 40 pM, 50 pM, 100 pM, 200 pM, 300 pM,and 1000 pM were injected into the in-line optofluidics. The process of each measurement was repeated as the aforementioned detection of 100 pM B[a]P solution. The in-line


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optofluidic was kept for 60s in order to allow all B[a]P molecules in the solution to bound to the immobilized antibodies completely. Figure 6(a) shows the wavelength shift of the in-line optofluidic with different concentrations of B[a]P solution. The wavelength shift of the in-line optofluidic indicates that the specific associations were occurring between the B[a]P molecule and antibody. Then the in-line optofluidic was washed by using pepsin solution (50 ppm in glycine HCl buffer; pH 2.0) to remove the B[a]P, followed by washing in ultrapure water for 80 s. The wavelength shift, as shown in Figure 6(a), was increased back to the baseline after each washing, indicating that the bound B[a]P molecule was released back into solution by the pepsin solution. Besides, for B[a]P solution concentrations lower than 50 pM (30 pM and 40 pM), the wavelength shift for the proposed biosensor remained almost constant, as shown in Figure 6(b)-(d). Therefore, the limit of detection achieved for the proposed in-line optofluidic is 50 pM. The relationship between the wavelength shift and the concentration of B[a]P solution is shown in Figure 6(e). The response of the in-line optofluidic maintains a good linear relationship between the wavelength shift and the concentration of B[a]P solution, and a sensitivity of 23pm/pM is achieved. In the experiments, the sensing response of the antibody immobilization, which is a large macromolecule, gives only ~2nm shift after washing, while the response of the B[a]P molecules with the concentration of 1000 pM is greater than 20nm, which can be attributed to the high refractive index of the B[a]P molecule: In general, the refractive index of the antibody is 1.45–1.55. Hence the refractive index of the B[a]P molecule (1.887) is much larger than that of the antibody [32], resulting in the large wavelength shift of the B[a]P molecule. In the proposed sensor, the response saturation of 2500 pM was achieved (see Supporting Information).


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Figure 6(a) The wavelength shift of the in-line optofluidic with different concentrations of B[a]P solution. The wavelength shift with concentrations of (b)50pM, (c) 40pM, and (d) 30pM. (e) The relationship between the wavelength shift and the concentration of B[a]P solution.

CONCLUSION In conclusion, an in-line fiber optofluidic sensor for the detection of B[a]P molecule in water is proposed and experimentally demonstrated. An air hole in the cladding of the HCF was fabricated as an in-line fiber optofluidic, into which the liquid sample was flowed through the inlet and outlet. Power leakage occurs at resonant wavelengths due to the guidance mechanism of the ARROW. The antibody was immobilized on the surface of the in-line fiber optofluidic, and the B[a]P molecule can be measured by interrogating the wavelength shift of the lossy dip due to the immunoreaction between the antibody and the target B[a]P. The experimental results show that sensitivities of up to 23 pm/pM, and the detection limit of 1.65 pM are achieved, establishing the in-line fiber optofluidic as a sensitive and versatile sensor. The proposed fiber sensor appears to have potential applications in research on environmental contamination, food safety, human health, etc.


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ASSOCIATED CONTENT Supporting Information Brief information in refractive index response, detection limit, saturation response, and regeneration of the sensor are supplied as Supporting Information.

AUTHOR INFORMATION Corresponding Author *Phone: +86-010-58887196. E-mail: [email protected].

Author Contributions R. G. conducted experiments. D. F. L. and M. Y. Z. carried out chemical synthesis. Z. M. Q. gave general guidance and advice.

ACKNOWLEDGMENT This work was supported by the National Basic Research Program of China (Nos. 2015CB352100), the National Natural Science Foundation of China (Nos. 61675203, 61377064, 61401432, 61501425, and 61601436), the Beijing Natural Science Foundation (4174108), and the Research Equipment Development Project of Chinese Academy of Sciences (No. YZ201508).

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Figure 1.(a) Cross-section of the HCPCF. (b) The schematic diagram of the proposed sensor. (c) Square micro-channel for the inlet and outlet.


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Figure 2 (a) The experimental setup of the proposed sensor. Inset shows the metal plate. The close-up view of the HCPCF. (b) Fluorescent image of the HCPCF filled with anti-mouse IgG -FITC in all hollow holes (c), and only one hollow hole (d).


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Figure 3 (a) Numerical simulation of the HCPCF without the liquid sample. (b) Schematic diagram of the cross-section of the liquid-infiltrated HCPCF. (c) Guiding mechanism of the liquid-infiltrated HCPCF. Numerical simulation of the HCPCF filled with liquid (d) at the anti-resonant wavelength (1525.00 nm), (e) at the resonant wavelength (1535.42 nm).


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Figure 4 (a) Transmission spectrum of the HCPCF with the liquid sample.(b) The transmission spectra of the sensor with the glutaricdialdehyde, antibody, BSA, and B[a]P. (c) The wavelength shift of the transmission spectra at various surface events after ultrapure water rinsing.


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Figure 5 (a)The time response of the in-line optofluidic. (b) The temperature response of the in-line optofluidic. (c) Comparison of response of in-line optofluidic to 100 pM B[a]P and other PAHs. (c) Wavelength shift for the water sample.


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Figure 6 (a)The wavelength shift of the in-line optofluidic with different concentrations of B[a]P solution. The wavelength shift with concentrations of (b)50pM, (c) 40pM, and (d) 30pM. (e) The relationship between the wavelength shift and the concentration of B[a]P solution.


Optofluidic immunosensor based on resonant wavelength shift of a

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