Abstract

We propose to use squeezed hollow-core photonic bandgap Bragg fibers for surface sensing applications. We demonstrate theoretically and confirm experimentally that squeezing a section of the Bragg fiber core increases overlap between the optical fields of the core guided modes and the modes bound to the sensing layer, thus, significantly enhancing their interaction via anticrossing phenomenon, which, in turn, enhances surface sensitivity of the fiber sensor. As a practical demonstration, we apply our fiber sensor to in situ monitoring of the dissolution dynamics of a sub-micron-thick polyvinyl butyral (PVB) film coated on the surface of the liquid-filled Bragg fiber core. Strong spectral shift is observed during the dissolution of the PVB film, and a surface spectral sensitivity of ~0.07nm/nm is achieved experimentally with aqueous analytes. The proposed fiber sensor offers a new sensing modality and opens new sensing applications for photonic bandgap fibers, such as real-time detection of binding and affinity, study of kinetics, etc. for a range of chemical and biological samples.

© 2016 Optical Society of America

1. Introduction

Precise detection of bio-layer thickness, trace amounts of biomolecules, as well as identification of bacteria pathogens, is important in numerous bio/chemical applications including diagnostics, genomics, proteomics, microbiology, and surface chemistry [1–4]. Optical biosensors are powerful analysis tools to address these applications by means of fluorescence-based techniques or label-free detection. Compared to fluorescence-based techniques, label-free biosensors forgo the need of laborious labeling of samples, and can detect target molecules in their natural forms by measuring changes in the analyte refractive index [5], optical absorption [6], or Raman scattering [7], in the presence of the biological agents. The label-free sensing modality allows in situ monitoring of surface kinetics, affinity, as well as other surface interactions.

Optical fibers constitute a very promising platform for the label-free biosensing applications, as they offer miniaturization, high degree of integration, and continuous quantitative and qualitative real-time analysis. Up to date, a wide range of optical fiber based biosensors have been proposed and developed. One common implementation for surface sensing is by using evanescent coupling of the total internal reflection (TIR) guided modes to the bio-layer. This sensing strategy has been applied in various configurations, such as tapered microfibers [8], U-shaped fiber [9], side-polished fiber [10], as well as long period fiber grating (LPFG) [11]. However, in such sensors, in order to ensure efficient overlap between the bio-layer and the evanescent field of the optical modes, certain fiber modifications have to be first carried out, such as stripping off a section of the fiber cladding, or tapering fibers into microfibers, which undermine the mechanical robustness of these fiber sensors and limit the interaction length between the optical modes and the analyte.

Another popular implementation of the fiber sensors for surface sensing is based on the phenomenon of surface plasmon resonance (SPR), which involves resonant excitation of the electromagnetic surface waves at the metal-dielectric interface. Being extremely sensitive to the refractive index changes of the surrounding medium, these sensors are particularly suitable for biosensing applications such as monitoring molecule interaction, DNA hybridization, and antigen-antibody reactions [12–14]. However, this technique requires a complicated fabrication process including fiber cladding stripping, and precisely controlled metal-layer deposition. Furthermore, the probing length of surface plasmon in the visible range is only on the order of 100nm [15], which limits its utility when detection of larger targets (such as bacteria with sizes of 0.5μm-10μm) is required.

In order to extend the probing depth of the surface wave to longer distances for macromolecular or bacteria detection, one can pursue biosensors operating at longer wavelengths (such as THz). Thus, in [16] the authors used a suspended core polyethylene THz fiber for detection of the E. coli bacteria. The highly porous structure of a suspended core fiber allows convenient access to the fiber core and to the evanescent part of the guided wave. It was demonstrated that selective binding of the E. coli bacteria to the surface of the fiber core can significantly influence the THz transmission properties of the suspended core fiber due to enhanced scattering and absorption losses caused by the bacteria. Changes in the fiber propagation loss can then be correlated to the concentration of the bacteria in the liquid analyte. Additionally, You et al. reported a polypropylene hollow-tube based on an anti-resonant reflecting hollow waveguidance (ARRHW) mechanism for sensing molecular overlayers adhered to the inner-core surface of tubes [17,18]. The strongly localized THz evanescent wave enables detection of the subwavelength-thick overlayers (several microns to tens of microns) attached on the inner surface of the hollow tube. Nevertheless, strong absorption of the THz waves in water limits its application in aqueous environment.

Using sensors based on the hollow-core photonic bandgap Bragg fibers operating in the visible-near infrared spectral range constitutes a viable alternative for surface sensing applications in aqueous analytes, as Bragg fibers inherently integrate optical detection with microfluidics, and water is relatively transparent in this spectral range. Due to the possibility of having large core (~1mm), the response time of the liquid-filled Bragg fibers could be reduced to under 1s, thus allowing for continuous on-line monitoring of the biological and chemical samples in a contained, highly integrated manner. Furthermore, large fiber core could be further functionalized with bio-recognition elements that can bind and progressively accumulate target biomolecules or bacteria, thus enabling detection specificity. In general, one distinguishes two types of Bragg fibers, namely, low-refractive index (RI)-contrast Bragg fibers and high-RI-contrast Bragg fibers. We note that low-RI-contrast Bragg fibers have certain advantages for liquid analyte sensing, compared to their high-RI-contrast counterparts. In fact, we have argued in [19, 20], that high-RI-contrast Bragg fibers show superior guidance in the gaseous cores (nc~1), which makes them highly suitable for gas sensing applications. Meanwhile, when these high-RI-contrast Bragg fibers are filled with aqueous solutions (nc~1.33), the TM bandgaps of the Bragg reflector tend to collapse at the light line of aqueous material due to the Brewster angle phenomenon, leading to high loss of the guided modes. In contrast, the low-RI-contrast Bragg fibers show large TM band gaps in the vicinity of the light line of an aqueous solution-filled fiber core, thus resulting in good guidance of all the polarizations. Therefore, the use of low-RI-contrast Bragg fibers is preferred in the case of the aqueous analytes.

In our previous work, we have demonstrated using hollow-core Bragg fibers for bulk refractive index sensing of liquid samples [21–23]. The operation principle of such sensors relies on the spectral modality, according to which variation in the analyte refractive index modifies the bandgap guidance of the fiber, leading to spectral shift in the fiber transmission spectra. In this paper, we demonstrate a novel application of the hollow-core Bragg fibers for surface sensing. The fiber used in this work is an all-polymer hollow-core low-RI-contrast Bragg fiber, which features a hollow core surrounded by a periodic sequence of the high/low refractive index multilayers, namely polystyrene (PS)/poly-methacrylate (PMMA) Bragg reflector. We then deposit a thin Polyvinyl butyral (PVB) analyte layer on the inner surface of the fiber core. The analyte layer supports its own guided modes with optical properties that are strongly affected by the layer refractive index and thickness. Due to the phenomenon of avoiding crossing in the vicinity of the phase matching wavelength between a particular pair of the core guided and the analyte layer bound modes, fiber transmission spectra (bandgap shape) can be significantly modified, thus allowing monitoring of changes in the analyte layer properties. It is, in fact, the hybridization of the core guided modes of a Bragg fiber with the analyte layer bound modes in the vicinity of their corresponding phase matching points that enables surface sensing modality of Bragg fibers. One of the design challenges when using Bragg fibers for surface sensing is a weak overlap between the modal fields of the core guided and the analyte layer bound modes. This is due to the fact that the fields of the core guided modes of a Bragg fiber decrease rapidly from the core center and towards the core inner surface where the analyte layer is located. In order to increase the field overlap and in order to promote hybridization between the core guided modes and the analyte layer bound modes, we propose squeezing of the Bragg fiber. Squeezing of the fiber decreases the fiber core size in one dimension, which, in turn, leads to the field enhancement and mode overlap in the direction of squeezing. As a demonstration of the surface sensing modality, we apply our sensor to monitoring the dissolution dynamics of the Polyvinyl butyral (PVB) film coated on the inner surface of the fiber core. Furthermore, we demonstrate that the surface sensitivity of the fiber sensor could be greatly enhanced by squeezing a section of the Bragg fiber, and thus, increasing the modal overlap between the core modes and the modes bound to a thin PVB film. Simulations based on the transfer matrix method are conducted to validate the experimental results. The proposed fiber sensor opens a new approach for a wide range of application, such as detection of molecular interaction, study of surface kinetics, as well as bacteria detection and other applications involving surfaces.

2. Theoretical analysis of using Bragg fiber for surface sensing applications

Bragg fibers are well known for their ability to confine light in the core filled with analyte having refractive indices lower than those of the fiber cladding materials. When a broadband light source is launched into the Bragg fiber core, only the light with frequencies within bandgap of the Bragg reflector will be confined and guided in the fiber core, while the light with frequencies outside of the bandgap will be irradiated out.

In order to theoretically study the surface sensing mechanism, we simulate the modal refractive indices and the propagation losses of the guided modes inside of the Bragg fibers by using the transfer matrix method (TMM) [24, 25], with an analyte layer deposited onto the core inner surface. In our simulations, we approximate the circular Bragg fibers as planar Bragg stacks with the same material composition and layer thickness. First, we consider a Bragg fiber with a bilayer featuring refractive indices of nh, nl, and having corresponding layer thicknesses of dh, dl. The core is assumed to be filled with water (nc = 1.33), and has a diameter of D. The number of bilayers in the fiber reflector is N. In Fig. 1, the green region represents an analyte layer with a thickness of da and refractive index of na.

 figure: Fig. 1

Fig. 1 Schematic of the Bragg fiber (a) and the corresponding planar Bragg waveguide (b) with the same material composition and layer thickness. The Bragg reflectors feature alternating polystyrene (PS) /poly-methacrylate (PMMA) layers. The green region represents the analyte layer.

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2.1 Influence of the analyte layer thickness on the Bragg fiber spectral features

In our first set of simulations of the squeezed Bragg fibers, we calculate the modal refractive indices and the radiation losses of the TM1 and TE1 modes guided in the water-filled fiber core that features a thin analyte layer with thickness ranging from 0nm to 200nm. The thicknesses of PMMA/PS bilayer are obtained from the scanning electro microscopy graphs of the cross section of the Bragg fiber used in our experiments (dPMMA = 210nm. dPS = 210nm), as shown below. The number of bilayers N in the Bragg reflector is 20. The refractive indices of the PMMA, PS, and water are (nPMMA = 1.49, nPS = 1.58, nwater = 1.33). The fiber core diameter is set to be D = 600μm. The refractive index of the analyte layer is set to be 1.46, which is a typical value for many bio-materials [26], as well as for the PVB layer used in our experiments.

Figures 2(a)-2(d) illustrate the simulated modal refractive indices and the radiation losses of the TM1 and TE1 modes guided in the Bragg fiber as a function of wavelength. We note that introduction of an analyte layer into the fiber core results in the anticrossing phenomenon when the fundamental mode of the Bragg fiber (closest to the light line of the core material) hybridizes with the modes guided predominately in the analyte layer. This phenomenon happens in the vicinity of the phase matching wavelength between the dispersion relations of the two modes, and it results in apparent discontinuities in the modal dispersion curves, corresponding to the anticrossings between the Bragg waveguide core modes and the so-called defect modes [27–29] guided in the analyte layer.

 figure: Fig. 2

Fig. 2 Simulated modal dispersion relations of TM1 (a) and TE1 (b) modes in the water-filled Bragg fiber featuring an analyte layer with thickness ranging from 0nm to 200nm. Corresponding radiation losses of the TM1 (c) and TE1 (d) modes. Dependence of the wavelength of avoided crossing of the TM1 (e) and TE1 (f) modes (obtained from (c) and (d)) as a function of the analyte layer thickness. Dependence of the wavelength of the lowest radiation loss of the TM1 (g) and TE1 (h) modes (obtained from (c) and (d)) as a function of the analyte layer thickness.

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A schematic view of anticrossing between a particular pair of a core guided mode and an analyte layer bound defect mode is illustrated in Fig. 3. The dispersion relation of a core guided mode of a circular Bragg fiber (without analyte layer) is shown as a blue dotted line. Similarly, the dispersion relation of an analyte layer bound mode (mode of a step-index slab waveguide) is shown as a purple dotted line. The wavelength of crossing between the two dispersion relations is called a phase match point between the two modes. Now, when plotting the corresponding modes of a Bragg fiber with analyte layer we observe anticrossing of the dispersion relations corresponding to the core guided and analyte bound mode in the vicinity of phase matching point. In the vicinity of the phase matching point (also wavelength of avoiding crossing), there is a resonant power transfer from the core mode of a Bragg fiber into the defect mode of an analyte layer. In turn, the defect mode in the analyte layer is phase matched with the radiation continuum of the fiber cladding, thus resulting in a significant increase of the fiber propagation losses in the vicinity of the wavelength of the avoiding crossing.

 figure: Fig. 3

Fig. 3 Schematic of the modal dispersion relations of the core guided (blue dotted line) and analyte layer bound modes (purple dotted line), as well as dispersion relations of the hybridized modes in a Bragg fiber with an analyte layer (green solid lines) near the point of phase matching (anticrossing wavelength).

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The wavelength of avoiding crossing depends strongly on the analyte layer thickness. When increasing the analyte layer thickness, the anticrossing wavelength shows a red shift. This is easy to rationalize by noting that dispersion relation of the fundamental mode of a slab waveguide (analyte layer) shows red shift when the layer thickness increases.

Moreover, we note that, upon introduction of the analyte layer, the original bandgap of the Bragg fiber (without the analyte layer) becomes fractured due to appearance of the loss peak near the wavelength of anticrossing. This new structure of the bandgap suggests two modalities for the detection of changes in the analyte layer thickness. The most direct modality (modality 1 in Figs. 2 (e) and 2(f)) is to track the wavelength of the avoided crossing that appears as a transmission dip inside of the original fiber bandgap (region of high transmission). From Figs. 2(e) and 2(f), one finds a nonlinear dependence of the wavelength shift with respect to the analyte layer thickness. However, in practice, this modality is difficult to realize as large core Bragg fibers are highly multimode, so there is no single anticrossing wavelength. Thus, in place of a single transmission dip one typically observes a change in the bandgap shape. This phenomenon is clearly observed in Figs. 2(c) and 2(d). Due to the avoided crossing, a transmission loss peak appears inside of a bandgap. This peak splits the bandgap into two smaller bandgaps. By tracking the shift in the wavelength of the lowest loss of a large bandgap (modality 2 in Figs. 2 (c) and 2(d)), one can indirectly track the changes in the position of the anticrossing wavelength (transmission dip in the fiber bandgap). We find in the following that modality 2 more relevant experimentally. In order to characterize the surface sensitivity of this modality, we thus plot the wavelength of the lowest propagation loss of the TM1 and TE1 modes as a function of the analyte layer thickness, as shown in Figs. 2(g) and 2(h). An approximately linear dependency with analyte layer thickness is found for both modes. The surface sensitivity of the spectral method can be then defined as Sλ=λc/da, where, λc is the wavelength shift when the analyte layer thickness is changed by da. According to this definition, the simulated surface sensitivity using modality 2 is found to be 0.046nm/nm and 0.04nm/nm for the TM1 and TE1 modes, respectively. We note that the simulated surface sensitivity is mostly valid for Bragg fiber operating in the single mode or a few modes regime. In our experiments, the hollow-core Bragg fiber features a relatively large core diameter (~600μm), and a relatively small length of 8cm, which allows thousands of modes propagating in the fiber core. As a result, many higher-order modes are excited when the light beam is launched into the fiber. These higher-order modes have somewhat different modal overlap with the analyte layer, as well as different anticrossing wavelengths and transmission bandgaps (see Fig. 3). Therefore, we expect that the surface sensitivity measured experimentally would be a result of combined contribution of many higher-order modes, rather than only the fundamental TE1 and TM1 modes. At the same time, we note that another effect of the multimode operation on the fiber transmission is in broadening of the transmission window and the overall increase in the propagation loss.

2.2 Effect of the sensor core size (degree of squeezing) on the surface sensitivity

Now, we study effect of the squeezed fiber core on the surface sensitivity of our sensor. When dealing with the squeezed Bragg fibers, we approximate them as planar Bragg stacks with the same material composition and layer thickness. When the fiber is squeezed, it can be considered that the distance between planar stacks D is reduced. In the following simulation, we keep the analyte layer thickness constant (100nm), and calculate the surface sensitivity of 10 possible even-symmetric TE and TM modes (closest to the light line) of the squeezed fiber when D is decreased from 600μm to 20μm. In Fig. 4, we plot the surface sensitivity of the 10 TE and TM modes as a function of the squeezed core size D. One can make several interesting observations from the simulation results. First, higher order modes show higher surface sensitivity compared to the fundamental modes, due to higher modal overlap with the analyte layer coated on the fiber inner surface. Second, when reducing the fiber core size, the surface sensitivity of the fiber sensor increases. This enhancement is especially pronounced when the squeezed core size becomes smaller than 100μm. This remarkable improvement of the surface sensitivity is attributed to an enhanced modal overlap between the core-guided modes and the analyte layer. This demonstrates that, indeed, the surface sensitivity of the fiber sensor can be enhanced dramatically by squeezing a section of the Bragg fiber core so as to increase the modal overlap with the analyte layer. From the simulation results, one can find that the largest sensitivity enhancement of the 10 TE and TM modes is ~52% when the fiber core size is reduced from 600μm to 100μm. Further, this enhancement could reach up to ~12-fold when the fiber core size is reduced to 20μm. We note that, in practice, our fiber sensor operates in a multimode regime. The extent of the surface sensitivity enhancement would be a result of combined contribution of the great number of optical modes propagating in the fiber core, with higher order modes being more sensitive to changes to the fiber core size.

 figure: Fig. 4

Fig. 4 Simulated surface sensitivity of 10 even symmetric TE (a) and TM (b) modes (closest to the light line) when the fiber core size D is reduced from 600μm to 20μm. The PVB layer thickness is kept constant and equal to 100nm in all the simulations.

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3. Experimental realization of the Bragg fiber sensor

3.1 Deposition of the PVB layer

In order to experimentally study the squeezed Bragg fiber sensors and assess their potential for surface sensing applications, we undertook the study of dissolution dynamics of a thin analyte film deposited on the fiber core inner surface. Particularly, a thin layer of Polyvinyl butyral (PVB) resin (Butvar® B-98, Sigma-Aldrich) is chosen as the analyte material in this demonstration because it is transparent in the visible spectral range, and because a PVB layer deposited on to the fiber core inner surface can be easily dissolved using ethyl alcohol without damaging the PS/PMMA reflector. A PVB layer fabricated using a solution-based method has a refractive index in the range of 1.45-1.47, which is close to that of many practical biomaterials.

In order to deposit a thin PVB film on the inner surface of the Bragg fiber, 1.5g PVB is first dissolved in 10 mL ethyl alcohol and mixed thoroughly. Then, the mixture is pumped into the Bragg fiber via applying a negative pressure. Subsequently, the 8cm long fiber section is fixed on the chuck of a spin coater (WS-400-6NPP, Laurel) and rotated at a speed of 3000RPM for 60s. The cross section of the obtained fiber with PVB layer inside is shown in Fig. 5 (a) and (b). From the SEM micrograph, the thickness of the PVB layer is estimated to be ~500nm, while the thicknesses of the individual PMMA and PS layers are both ~210nm.

 figure: Fig. 5

Fig. 5 (a) Cross section of the hollow-core PMMA/PS Bragg fiber with a PVB layer under optical microscope. (b) Cross section of the Bragg reflector taken by a scanning electron microscope (SEM), featuring the alternating polystyrene (PS) and poly-methacrylate (PMMA) layers with individual thickness of ~210nm, as well as a ~500nm-thick PVB layer.

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3.2 Experimental setup

Experimentally, we use two opto-fluidic blocks to integrate the Bragg fiber sensor into the setup, as shown in Fig. 6. The two opto-fluidic blocks enable optical coupling and flow of the target analytes, which are detailed in our prior work [21,30]. After pumping liquid analyte into the hollow core Bragg fiber, a beam from a broadband supercontinuum source (500nm-1700nm) is launched into one end of the fiber core using an objective ( × 10). The transmitted spectrum is registered by a grating based monochromator (Oriel, 1/8m cornerstone, Newport). In order to enhance surface sensitivity of the Bragg fiber sensor, we increase the modal overlap between the fiber core modes and the analyte film by squeezing the fiber using a metallic rod fixed on a micro-positioning stage. The fiber is squeezed in increments of 50μm.

 figure: Fig. 6

Fig. 6 Experimental setup for the optical characterization of the surface sensitivity of the squeezed Bragg fiber sensor. The 8cm-long Bragg fiber with a thin PVB coated on the inner surface is integrated into the set-up using two opto-fluidic blocks. A broadband supercontinuum beam is launched into one end of the liquid-core Bragg fiber using an objective, while the output spectrum of the fiber sensor is registered by a monochromator. The metallic rod (diameter: 2.54cm) is fixed on a micro-positioning stage and is used to squeeze a small section of the fiber in increments of 50μm. Inset: ~8cm Bragg fiber used in this experiment (top), squeezing sketch of our setup (middle), as well as the squeezed section of the Bragg fiber (bottom).

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3.3 Characterization of dissolution dynamics of the PVB analyte layer

First, we characterize the surface sensitivity of the Bragg fiber sensor without squeezing. The fiber features a PVB analyte layer (with a thickness of ~500nm) at the start. We note that since a grating based spectrometer is used to register the transmission spectra, each scan of the spectral region of interest (550-750nm) is relatively large, and it takes ~60 seconds. When introducing ethyl alcohol into fiber, the analyte layer starts dissolving immediately and the Bragg fiber spectra change continuously. Therefore, before pumping ethyl alcohol into the fiber (in order to dissolve the PVB layer), we first register the reference transmission spectrum using chemically inert 16%wt NaCl solution in water. This NaCl solution is chosen because it has the same refractive index (n~1.36) as ethyl alcohol and thereby the fiber has the same bandgap position as the one filled with ethyl alcohol. After acquiring the reference spectrum, the ethyl alcohol is introduced into the fiber core in order to dissolve a thin PBV layer deposited onto the fiber core inner surface. The time varying transmission spectra are then acquired every minute. As shown in Fig. 7, with the dissolution of a PVB layer, the transmission spectra have an obvious blue shift towards shorter wavelengths. The direction of the spectral shift is consistent with that found in numerical simulations. Moreover, as the PVB film (RI~1.46) dissolves in ethyl alcohol (RI~1.36), we expect that the bulk refractive index of the fiber core should gradually increase, which induces an additional blue spectral shift [19, 21]. This assumption is well verified by noting that when we purge the fiber with pure ethyl alcohol after the PVB layer is dissolved, a red shift of ~5nm is observed, which is due to the decrease of the refractive index of pure ethyl alcohol compared to that of the PVB-alcohol mixture. Therefore, based on the observed results, we can make the following statement, the surface sensitivity of a Bragg fiber is made of two contributions. The first contribution is due to a true surface sensing modality owing to thickness variation of an analyte layer attached to the inner surface of the fiber core, while the second bulk contribution is due to changes of the bulk refractive index of the fiber core due to dissolution of the analyte layer in the liquid material of a fiber core (assuming that the core liquid is not continuously pumped through the fiber core).

 figure: Fig. 7

Fig. 7 Experimental transmission spectra of the Bragg fiber during the dissolution process of a thin PVB layer (thickness~500nm) coated on the inner surface of the fiber core. A blue shift is observed with dissolution of the PVB film. After the PVB layer is dissolved, the fiber is purged with fresh ethyl alcohol, and a red shift of ~5nm is observed, which is attributed to the fact that the RI of ethyl alcohol is somewhat lower than that of the PVB-alcohol mixture.

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Finally, we refill the fiber used in the prior experiment with 16%wt NaCl solution in distilled water, and acquire their spectra after a complete removal of the PVB layer. Comparison with the initial transmission peak positions that are measured before the dissolution process allows us to calculate the surface sensitivity of the Bragg fiber. In Fig. 8, we plot the spectra of the Bragg fiber filled with 16%wt NaCl in water before and after the PVB film dissolution. We find that presence of a ~500nm thick PVB layer leads to 23nm red shift in the fiber transmission spectrum, compared to the spectrum of the fiber without a PVB layer. The corresponding surface sensitivity of our sensor to changes in the layer thickness is then estimated to be 0.046nm/nm. For completeness, we also present in Fig. 8(a) the transmission spectra of a Bragg fiber with and without PVB layer when the fiber core is filled with distilled water. In this case, we also observe a ~26nm shift of the transmission peak, thus resulting in a surface sensitivity of 0.052nm/nm. The experimentally demonstrated surface sensitivity is comparable to that found in the numerical simulations. Additionally, from the time dependent date of Fig. 7, we can estimate that the dissolution rate the PVB layer is ~2nm/s. We note that in the spectra of a Bragg fiber with a PVB layer, a pronounced transmission dip at the left edge of the fiber transmission window is well observed both in the case of water and 16%wt NaCl solution (see Figs. 8(a) and 8(b)). The presence of this transmission dip is probably attributed to anticrossing between the fiber core modes and the modes confined to the analyte PVB layer, as we have observed in the simulated spectra in Section 2.

 figure: Fig. 8

Fig. 8 (a) Experimental transmission spectra of the ~8cm-long Bragg fiber filled with distilled water (a), (c), and 16%wt NaCl in water (b), (d), before and after the PVB film dissolution. The ~500nm PVB film induces a red shift of 26nm and 23nm in the transmission peak, respectively, when compared to transmission of the same fiber without a PVB layer.

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4. Enhancement of the surface sensitivity by squeezing the Bragg fiber

Due to large core of the Bragg fiber (significantly larger than the wavelength of light and the analyte layer thickness), only a small fraction of power of the core-guided mode is present in the vicinity of the inner surface of the fiber core, thus leading to a poor modal overlap with the PVB analyte layer, therefore, only a moderate surface sensitivity is observed experimentally. In order to enhance the surface sensitivity, numerical simulations of Section 2 suggest that one has to reduce the fiber core size. To verify this experimentally, we repeated the surface sensitivity characterization experiment when a section of the Bragg fiber is squeezed, thus, effectively reducing the fiber core size. Advantage of fiber squeezing versus drawing fiber of a smaller diameter is that in the case of fiber squeezing one can explore the tradeoff between sensitivity enhancement and sensor response time increase observed for smaller fiber core sizes. More importantly, such a fiber-squeezing technique can also serves as a flexible and independent ‘knob’ to adjust the surface sensitivity and the response time, to suit particular needs for different applications.

The experimental setup for the squeezed fiber experiment is shown in Fig. 6. The metallic rod with a diameter of 2.54cm is fixed on the micro-positioning stage, which enables squeezing a one-centimeter-long section of the fiber (as shown in Fig. 6) in increments of 50μm. We choose the same Bragg fiber, as well as the same recipe for the PVB film deposition in order to guarantee similar transmission properties of the Bragg fiber and similar PVB layer thicknesses as in the case of experiments with non-squeezed fibers detailed in Section 3.

In Fig. 9, we plot the transmission spectra of the Bragg fiber filled with distilled water when a section of the fiber is squeezed by displacements ranging from 0μm to 500μm. The fiber has a PVB layer with a thickness of ~500nm. We can see that although the transmission amplitude of the fiber is decreased with the increase of the squeezing displacement, the transmission peak positions are well conserved during the squeezing process.

 figure: Fig. 9

Fig. 9 Experimental transmission spectra of the ~8cm long Bragg fiber filled with distilled water before PVB layer dissolution, when a section of fiber is squeezed by displacements ranging from 0μm to 500μm using a 2.54cm diameter cylinder. The fiber has a PVB layer with thickness of 500nm.

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Then, we registered the time varying spectra of the Bragg fiber during the dissolution of a PVB layer using ethyl alcohol when a section of the fiber is squeezed. Two squeezing displacements of 250μm and 500μm are used in our preliminary experiments. In Fig. 10, we present the spectra of the Bragg fiber filled with distilled water and 16%wt NaCl solution in water before and after the complete dissolution of the PVB layer with a goal of characterizing the overall surface sensitivities and their enhancement due to fiber squeezing. We remind that the reader that refractive index of the 16%wt NaCl solutions in water is the same as that of ethyl alcohol that we use for dissolution of the PVB layer. At the same time, PVB is not dissolvable in NaCl solution, which motivates us to use such a solution for reference measurements before and after PVB layer dissolution. Measurement with pure distilled water is presented in lieu of the potential applications of our sensor with aqueous analytes and bio-functional layers.

 figure: Fig. 10

Fig. 10 Experimental transmission spectra of the ~8cm long Bragg fiber filled with distilled water and 16%wt NaCl solution in water before and after the PVB film dissolution. The fiber is squeezed by the depths of 250μm and 500μm using a 2.54cm diameter cylinder. In the case of a 250μm-squeezed fiber (D~350μm), a ~500nm PVB film induces a red shift of 30nm (filled with pure water), and 26nm (filled with 16%wt NaCl solution). In the case of a 500μm-squeezed fiber (D~100μm), the 500nm PVB film induces a red shift of 35nm (filled with pure water), and 29nm (filled with 16%wt NaCl solution).

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By comparing the transmission peak positions of the fiber filled with distilled water and 16%wt NaCl solution before and after the PVB film is fully dissolved, we can calculate the surface sensitivity of the Bragg fiber at those two different degrees of squeezing. As shown in Fig. 11, the surface sensitivities of a squeezed fiber are enhanced by ~35% and ~30% when the fiber is filled with distilled water and 16%wt NaCl solution, respectively, compared to that of a non-squeezed fiber. The extent of sensitivity enhancement shows reasonable agreement with our simulation results. The mechanism of the sensitivity enhancement is due to increased overlap between the core modes and the PVB analyte layer when the fiber is squeezed.

 figure: Fig. 11

Fig. 11 Surface sensitivities of the squeezed hollow core Bragg fiber sensor filled with distilled water and 16%wt NaCl solution for various degrees of squeezing. The circular fiber core has a diameter of ~600μm. Inset: Schematic of the squeezed Bragg fiber with core size of D.

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5. Discussion

In this paper, we propose to use squeezed hollow-core Bragg fibers for surface sensing applications. As a demonstration, we apply the fiber sensor to monitor the thickness variation of a thin analyte layer deposited on the inner surface of the fiber core. When a thin film is introduced onto the fiber inner surface, anticrossings between core guided modes and modes confined to the analyte layer are observed both theoretically and experimentally. In the vicinity of anticrossings, the optical energy is resonantly transferred from the fiber core modes to the analyte layer modes, followed by a strong increase in the fiber propagation loss. The wavelength of modal anticrossing depends strongly on the analyte thickness, thus allowing monitoring of changes in the analyte layer thickness. The presence of the attenuation peaks induced by anticrossing phenomenon deforms the shape of the original bandgap of the Bragg fiber, and results in changes in the shape of a transmission spectrum. Experimentally, it is easier to monitor spectral shift of the point of the lowest transmission loss than the anticrossing wavelength (transmission dip). This spectral modality for the monitoring of the analyte layer optical properties is suitable for a wide range of surface sensing applications.

Due to large size of the Bragg fiber core (which is much larger than the thickness of the analyte layer), only a poor modal overlap between the core guided modes and the analyte layer modes is achieved, thereby leading to moderate surface sensitivity. In order to enhance surface sensitivity, we explore squeezing a section of the Bragg fiber, which enhances guided modes/analyte layer overlap. Numerical simulations show that, squeezing the hollow-core Bragg fiber could significantly enhance the surface sensitivity. This enhancement is especially pronounced when the squeezed core size becomes smaller than 100μm. However, due to moderate mechanical robustness and chemical resistivity of the all-polymer Bragg fiber when it is filled with ethyl alcohol, the fiber breaks or cracks if squeezed by a displacement of more than 500μm (the resultant core size ~100μm). As a result, only a moderate sensitivity enhancement (~35%) is achieved in our experimental demonstration when the core size is reduced from 600μm to 100μm. Despite the limited degrees of squeezing, we can still clearly observe enhancement of the surface sensitivity when using squeezing of the fiber core. In our future work, we will study the ways of increasing mechanical robustness of the polymer Bragg fibers to squeezing in order to further enhance surface sensitivity.

6. Conclusion

In summary, in this work, we study both theoretically and experimentally the use of hollow-core Bragg fibers for surface sensing applications. The fiber sensor operates using a spectral sensing modality to monitor changes in the thickness of an analyte layer deposited on the inner surface of the Bragg fiber core. As a practical demonstration, we have applied this sensor to monitor the dissolution dynamics of a thin PVB layer coated on the fiber inner surface. A surface sensitivity of 0.052nm/nm is experimentally achieved with aqueous analyte. Moreover, we have also experimentally demonstrated that squeezing of a section of the Bragg fiber, increases the overlap between the core-guided modes and the analyte layer, thus enhancing the surface sensitivity of the fiber sensor. The highest surface sensitivity achieved in our work is 0.07nm/nm with the squeezed fiber core size of 100μm. The experimental observations are validated by theoretical simulation based on the transfer matrix method. The proposed liquid-core Bragg fiber sensor is applicable to a wide range of surface sensing applications including molecular recognition, bacteria detection, and monitoring of the biolayer thickness variation and others, with the advantages of simplicity in sensing mechanism, short response time, ease of fabrication, and relatively high sensitivity.

References and links

1. M. Zourob, S. Elwary, and A. Turner, Principles of Bacterial Detection: Biosensors, Recognition Receptors and Microsystems (Springer, 2008).

2. G. Gauglitz, “Direct optical sensors: principles and selected applications,” Anal. Bioanal. Chem. 381(1), 141–155 (2005). [CrossRef]   [PubMed]  

3. X. D. Wang and O. S. Wolfbeis, “Fiber-optic chemical sensors and biosensors (2008-2012),” Anal. Chem. 85(2), 487–508 (2013). [CrossRef]   [PubMed]  

4. X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008). [CrossRef]   [PubMed]  

5. E. Chow, A. Grot, L. W. Mirkarimi, M. Sigalas, and G. Girolami, “Ultracompact biochemical sensor built with two-dimensional photonic crystal microcavity,” Opt. Lett. 29(10), 1093–1095 (2004). [CrossRef]   [PubMed]  

6. J. B. Jensen, L. H. Pedersen, P. E. Hoiby, L. B. Nielsen, T. P. Hansen, J. R. Folkenberg, J. Riishede, D. Noordegraaf, K. Nielsen, A. Carlsen, and A. Bjarklev, “Photonic crystal fiber based evanescent-wave sensor for detection of biomolecules in aqueous solutions,” Opt. Lett. 29(17), 1974–1976 (2004). [CrossRef]   [PubMed]  

7. T. W. Koo, S. Chan, and A. A. Berlin, “Single-molecule detection of biomolecules by surface-enhanced coherent anti-Stokes Raman scattering,” Opt. Lett. 30(9), 1024–1026 (2005). [CrossRef]   [PubMed]  

8. W. B. Ji, S. C. Tjin, B. Lin, and C. L. Ng, “Highly sensitive refractive index sensor based on adiabatically tapered microfiber long period gratings,” Sensors (Basel) 13(10), 14055–14063 (2013). [CrossRef]   [PubMed]  

9. A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011). [CrossRef]  

10. A. Dudus, R. Blue, and D. Uttamchandani, “Comparative study of microfiber and side-polished optical fiber sensors for refractometry in microfluidics,” IEEE Sens. J. 13(5), 1594–1601 (2013). [CrossRef]  

11. M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000). [CrossRef]   [PubMed]  

12. A. K. Sharma, R. Jha, and B. D. Gupta, “Fiber-optic sensors based on surface plasmon resonance: a comprehensive review,” IEEE Sens. J. 7(8), 1118–1129 (2007). [CrossRef]  

13. T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013). [CrossRef]   [PubMed]  

14. J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009). [CrossRef]   [PubMed]  

15. A. Hassani and M. Skorobogatiy, “Photonic crystal fiber-based plasmonic sensors for the detection of biolayer thickness,” J. Opt. Soc. Am. B 26(8), 1550–1557 (2009). [CrossRef]  

16. A. Mazhorova, A. Markov, A. Ng, R. Chinnappan, O. Skorobogata, M. Zourob, and M. Skorobogatiy, “Label-free bacteria detection using evanescent mode of a suspended core terahertz fiber,” Opt. Express 20(5), 5344–5355 (2012). [CrossRef]   [PubMed]  

17. B. You, J. Y. Lu, J. H. Liou, C. P. Yu, H. Z. Chen, T. A. Liu, and J. L. Peng, “Subwavelength film sensing based on terahertz anti-resonant reflecting hollow waveguides,” Opt. Express 18(18), 19353–19360 (2010). [CrossRef]   [PubMed]  

18. C. H. Lai, B. You, J. Y. Lu, T. A. Liu, J. L. Peng, C. K. Sun, and H. C. Chang, “Modal characteristics of antiresonant reflecting pipe waveguides for terahertz waveguiding,” Opt. Express 18(1), 309–322 (2010). [CrossRef]   [PubMed]  

19. H. Qu and M. Skorobogatiy, “Liquid-core low-refractive-index-contrast Bragg fiber sensor,” Appl. Phys. Lett. 98(20), 201114 (2011). [CrossRef]  

20. M. Skorobogatiy, “Efficient antiguiding of TE and TM polarizations in low-index core waveguides without the need for an omnidirectional reflector,” Opt. Lett. 30(22), 2991–2993 (2005). [CrossRef]   [PubMed]  

21. J. Li, H. Qu, and M. Skorobogatiy, “Simultaneous monitoring the real and imaginary parts of the analyte refractive index using liquid-core photonic bandgap Bragg fibers,” Opt. Express 23(18), 22963–22976 (2015). [CrossRef]   [PubMed]  

22. H. Qu and M. Skorobogatiy, “Resonant bio- and chemical sensors using low-refractive-index-contrast liquid-core Bragg fibers,” Sens. Actuators B Chem. 161(1), 261–268 (2012). [CrossRef]  

23. H. Qu, B. Ung, M. Roze, and M. Skorobogatiy, “All photonic bandgap fiber spectroscopic system for detection of refractive index changes in aqueous analytes,” Sens. Actuators B Chem. 161(1), 235–243 (2012). [CrossRef]  

24. M. Skorobogatiy, Nanostructured and Subwavelength Waveguide: Fundamentals and Applications (Wiley, 2012).

25. M. Skorobogatiy and J. Yang, Fundamentals of Photonic Crystal Guiding (Cambridge University Press, 2009).

26. L. Rindorf, J. B. Jensen, M. Dufva, L. H. Pedersen, P. E. Høiby, and O. Bang, “Photonic crystal fiber long-period gratings for biochemical sensing,” Opt. Express 14(18), 8224–8231 (2006). [CrossRef]   [PubMed]  

27. J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008). [CrossRef]  

28. A. Argyros, I. M. Bassett, M. A. Eijkelenborg, and M. C. J. Large, “Analysis of ring-structured Bragg fibres for single TE mode guidance,” Opt. Express 12(12), 2688–2698 (2004). [CrossRef]   [PubMed]  

29. J. West, C. Smith, N. Borrelli, D. Allan, and K. Koch, “Surface modes in air-core photonic band-gap fibers,” Opt. Express 12(8), 1485–1496 (2004). [CrossRef]   [PubMed]  

30. J. Li, H. Qu, and M. Skorobogatiy, “Detection of analyte refractive index and concentration using liquid-core photonic Bragg fibers,” Proc. SPIE 9702, 97020R (2016). [CrossRef]  

References

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  • |

  1. M. Zourob, S. Elwary, and A. Turner, Principles of Bacterial Detection: Biosensors, Recognition Receptors and Microsystems (Springer, 2008).
  2. G. Gauglitz, “Direct optical sensors: principles and selected applications,” Anal. Bioanal. Chem. 381(1), 141–155 (2005).
    [Crossref] [PubMed]
  3. X. D. Wang and O. S. Wolfbeis, “Fiber-optic chemical sensors and biosensors (2008-2012),” Anal. Chem. 85(2), 487–508 (2013).
    [Crossref] [PubMed]
  4. X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
    [Crossref] [PubMed]
  5. E. Chow, A. Grot, L. W. Mirkarimi, M. Sigalas, and G. Girolami, “Ultracompact biochemical sensor built with two-dimensional photonic crystal microcavity,” Opt. Lett. 29(10), 1093–1095 (2004).
    [Crossref] [PubMed]
  6. J. B. Jensen, L. H. Pedersen, P. E. Hoiby, L. B. Nielsen, T. P. Hansen, J. R. Folkenberg, J. Riishede, D. Noordegraaf, K. Nielsen, A. Carlsen, and A. Bjarklev, “Photonic crystal fiber based evanescent-wave sensor for detection of biomolecules in aqueous solutions,” Opt. Lett. 29(17), 1974–1976 (2004).
    [Crossref] [PubMed]
  7. T. W. Koo, S. Chan, and A. A. Berlin, “Single-molecule detection of biomolecules by surface-enhanced coherent anti-Stokes Raman scattering,” Opt. Lett. 30(9), 1024–1026 (2005).
    [Crossref] [PubMed]
  8. W. B. Ji, S. C. Tjin, B. Lin, and C. L. Ng, “Highly sensitive refractive index sensor based on adiabatically tapered microfiber long period gratings,” Sensors (Basel) 13(10), 14055–14063 (2013).
    [Crossref] [PubMed]
  9. A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
    [Crossref]
  10. A. Dudus, R. Blue, and D. Uttamchandani, “Comparative study of microfiber and side-polished optical fiber sensors for refractometry in microfluidics,” IEEE Sens. J. 13(5), 1594–1601 (2013).
    [Crossref]
  11. M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
    [Crossref] [PubMed]
  12. A. K. Sharma, R. Jha, and B. D. Gupta, “Fiber-optic sensors based on surface plasmon resonance: a comprehensive review,” IEEE Sens. J. 7(8), 1118–1129 (2007).
    [Crossref]
  13. T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013).
    [Crossref] [PubMed]
  14. J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
    [Crossref] [PubMed]
  15. A. Hassani and M. Skorobogatiy, “Photonic crystal fiber-based plasmonic sensors for the detection of biolayer thickness,” J. Opt. Soc. Am. B 26(8), 1550–1557 (2009).
    [Crossref]
  16. A. Mazhorova, A. Markov, A. Ng, R. Chinnappan, O. Skorobogata, M. Zourob, and M. Skorobogatiy, “Label-free bacteria detection using evanescent mode of a suspended core terahertz fiber,” Opt. Express 20(5), 5344–5355 (2012).
    [Crossref] [PubMed]
  17. B. You, J. Y. Lu, J. H. Liou, C. P. Yu, H. Z. Chen, T. A. Liu, and J. L. Peng, “Subwavelength film sensing based on terahertz anti-resonant reflecting hollow waveguides,” Opt. Express 18(18), 19353–19360 (2010).
    [Crossref] [PubMed]
  18. C. H. Lai, B. You, J. Y. Lu, T. A. Liu, J. L. Peng, C. K. Sun, and H. C. Chang, “Modal characteristics of antiresonant reflecting pipe waveguides for terahertz waveguiding,” Opt. Express 18(1), 309–322 (2010).
    [Crossref] [PubMed]
  19. H. Qu and M. Skorobogatiy, “Liquid-core low-refractive-index-contrast Bragg fiber sensor,” Appl. Phys. Lett. 98(20), 201114 (2011).
    [Crossref]
  20. M. Skorobogatiy, “Efficient antiguiding of TE and TM polarizations in low-index core waveguides without the need for an omnidirectional reflector,” Opt. Lett. 30(22), 2991–2993 (2005).
    [Crossref] [PubMed]
  21. J. Li, H. Qu, and M. Skorobogatiy, “Simultaneous monitoring the real and imaginary parts of the analyte refractive index using liquid-core photonic bandgap Bragg fibers,” Opt. Express 23(18), 22963–22976 (2015).
    [Crossref] [PubMed]
  22. H. Qu and M. Skorobogatiy, “Resonant bio- and chemical sensors using low-refractive-index-contrast liquid-core Bragg fibers,” Sens. Actuators B Chem. 161(1), 261–268 (2012).
    [Crossref]
  23. H. Qu, B. Ung, M. Roze, and M. Skorobogatiy, “All photonic bandgap fiber spectroscopic system for detection of refractive index changes in aqueous analytes,” Sens. Actuators B Chem. 161(1), 235–243 (2012).
    [Crossref]
  24. M. Skorobogatiy, Nanostructured and Subwavelength Waveguide: Fundamentals and Applications (Wiley, 2012).
  25. M. Skorobogatiy and J. Yang, Fundamentals of Photonic Crystal Guiding (Cambridge University Press, 2009).
  26. L. Rindorf, J. B. Jensen, M. Dufva, L. H. Pedersen, P. E. Høiby, and O. Bang, “Photonic crystal fiber long-period gratings for biochemical sensing,” Opt. Express 14(18), 8224–8231 (2006).
    [Crossref] [PubMed]
  27. J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
    [Crossref]
  28. A. Argyros, I. M. Bassett, M. A. Eijkelenborg, and M. C. J. Large, “Analysis of ring-structured Bragg fibres for single TE mode guidance,” Opt. Express 12(12), 2688–2698 (2004).
    [Crossref] [PubMed]
  29. J. West, C. Smith, N. Borrelli, D. Allan, and K. Koch, “Surface modes in air-core photonic band-gap fibers,” Opt. Express 12(8), 1485–1496 (2004).
    [Crossref] [PubMed]
  30. J. Li, H. Qu, and M. Skorobogatiy, “Detection of analyte refractive index and concentration using liquid-core photonic Bragg fibers,” Proc. SPIE 9702, 97020R (2016).
    [Crossref]

2016 (1)

J. Li, H. Qu, and M. Skorobogatiy, “Detection of analyte refractive index and concentration using liquid-core photonic Bragg fibers,” Proc. SPIE 9702, 97020R (2016).
[Crossref]

2015 (1)

2013 (4)

X. D. Wang and O. S. Wolfbeis, “Fiber-optic chemical sensors and biosensors (2008-2012),” Anal. Chem. 85(2), 487–508 (2013).
[Crossref] [PubMed]

W. B. Ji, S. C. Tjin, B. Lin, and C. L. Ng, “Highly sensitive refractive index sensor based on adiabatically tapered microfiber long period gratings,” Sensors (Basel) 13(10), 14055–14063 (2013).
[Crossref] [PubMed]

A. Dudus, R. Blue, and D. Uttamchandani, “Comparative study of microfiber and side-polished optical fiber sensors for refractometry in microfluidics,” IEEE Sens. J. 13(5), 1594–1601 (2013).
[Crossref]

T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013).
[Crossref] [PubMed]

2012 (3)

A. Mazhorova, A. Markov, A. Ng, R. Chinnappan, O. Skorobogata, M. Zourob, and M. Skorobogatiy, “Label-free bacteria detection using evanescent mode of a suspended core terahertz fiber,” Opt. Express 20(5), 5344–5355 (2012).
[Crossref] [PubMed]

H. Qu and M. Skorobogatiy, “Resonant bio- and chemical sensors using low-refractive-index-contrast liquid-core Bragg fibers,” Sens. Actuators B Chem. 161(1), 261–268 (2012).
[Crossref]

H. Qu, B. Ung, M. Roze, and M. Skorobogatiy, “All photonic bandgap fiber spectroscopic system for detection of refractive index changes in aqueous analytes,” Sens. Actuators B Chem. 161(1), 235–243 (2012).
[Crossref]

2011 (2)

H. Qu and M. Skorobogatiy, “Liquid-core low-refractive-index-contrast Bragg fiber sensor,” Appl. Phys. Lett. 98(20), 201114 (2011).
[Crossref]

A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
[Crossref]

2010 (2)

2009 (2)

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

A. Hassani and M. Skorobogatiy, “Photonic crystal fiber-based plasmonic sensors for the detection of biolayer thickness,” J. Opt. Soc. Am. B 26(8), 1550–1557 (2009).
[Crossref]

2008 (2)

X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
[Crossref] [PubMed]

J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
[Crossref]

2007 (1)

A. K. Sharma, R. Jha, and B. D. Gupta, “Fiber-optic sensors based on surface plasmon resonance: a comprehensive review,” IEEE Sens. J. 7(8), 1118–1129 (2007).
[Crossref]

2006 (1)

2005 (3)

2004 (4)

2000 (1)

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

Allan, D.

Argyros, A.

Bang, O.

Bassett, I. M.

Bentley, W. E.

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

Berlin, A. A.

Bjarklev, A.

Blue, R.

A. Dudus, R. Blue, and D. Uttamchandani, “Comparative study of microfiber and side-polished optical fiber sensors for refractometry in microfluidics,” IEEE Sens. J. 13(5), 1594–1601 (2013).
[Crossref]

Bock, W. J.

A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
[Crossref]

Borrelli, N.

Campopiano, S.

A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
[Crossref]

Carlsen, A.

Chan, S.

Chang, H. C.

C. H. Lai, B. You, J. Y. Lu, T. A. Liu, J. L. Peng, C. K. Sun, and H. C. Chang, “Modal characteristics of antiresonant reflecting pipe waveguides for terahertz waveguiding,” Opt. Express 18(1), 309–322 (2010).
[Crossref] [PubMed]

J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
[Crossref]

Chang, T. C.

T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013).
[Crossref] [PubMed]

Chau, L. K.

T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013).
[Crossref] [PubMed]

Chen, H. W.

J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
[Crossref]

Chen, H. Z.

Chinnappan, R.

Chow, E.

Cusano, A.

A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
[Crossref]

Cutolo, A.

A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
[Crossref]

Davis, C. C.

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

DeLisa, M. P.

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

Delport, F.

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

Dudus, A.

A. Dudus, R. Blue, and D. Uttamchandani, “Comparative study of microfiber and side-polished optical fiber sensors for refractometry in microfluidics,” IEEE Sens. J. 13(5), 1594–1601 (2013).
[Crossref]

Dufva, M.

Eijkelenborg, M. A.

Fan, X.

X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
[Crossref] [PubMed]

Folkenberg, J. R.

Gauglitz, G.

G. Gauglitz, “Direct optical sensors: principles and selected applications,” Anal. Bioanal. Chem. 381(1), 141–155 (2005).
[Crossref] [PubMed]

Girolami, G.

Grot, A.

Gupta, B. D.

A. K. Sharma, R. Jha, and B. D. Gupta, “Fiber-optic sensors based on surface plasmon resonance: a comprehensive review,” IEEE Sens. J. 7(8), 1118–1129 (2007).
[Crossref]

Hansen, T. P.

Hassani, A.

Hoiby, P. E.

Høiby, P. E.

Hsieh, W. H.

T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013).
[Crossref] [PubMed]

Iadicicco, A.

A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
[Crossref]

Jans, K.

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

Janssen, K. P.

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

Jensen, J. B.

Jha, R.

A. K. Sharma, R. Jha, and B. D. Gupta, “Fiber-optic sensors based on surface plasmon resonance: a comprehensive review,” IEEE Sens. J. 7(8), 1118–1129 (2007).
[Crossref]

Ji, W. B.

W. B. Ji, S. C. Tjin, B. Lin, and C. L. Ng, “Highly sensitive refractive index sensor based on adiabatically tapered microfiber long period gratings,” Sensors (Basel) 13(10), 14055–14063 (2013).
[Crossref] [PubMed]

Koch, K.

Koo, T. W.

Lai, C. H.

Lammertyn, J.

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

Large, M. C. J.

Li, J.

J. Li, H. Qu, and M. Skorobogatiy, “Detection of analyte refractive index and concentration using liquid-core photonic Bragg fibers,” Proc. SPIE 9702, 97020R (2016).
[Crossref]

J. Li, H. Qu, and M. Skorobogatiy, “Simultaneous monitoring the real and imaginary parts of the analyte refractive index using liquid-core photonic bandgap Bragg fibers,” Opt. Express 23(18), 22963–22976 (2015).
[Crossref] [PubMed]

Li, Y. T.

J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
[Crossref]

Lin, B.

W. B. Ji, S. C. Tjin, B. Lin, and C. L. Ng, “Highly sensitive refractive index sensor based on adiabatically tapered microfiber long period gratings,” Sensors (Basel) 13(10), 14055–14063 (2013).
[Crossref] [PubMed]

Liou, J. H.

Liu, T. A.

Lu, J. Y.

Maes, G.

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

Markov, A.

Mazhorova, A.

Mirkarimi, L. W.

Ng, A.

Ng, C. L.

W. B. Ji, S. C. Tjin, B. Lin, and C. L. Ng, “Highly sensitive refractive index sensor based on adiabatically tapered microfiber long period gratings,” Sensors (Basel) 13(10), 14055–14063 (2013).
[Crossref] [PubMed]

Nielsen, K.

Nielsen, L. B.

Noordegraaf, D.

Paladino, D.

A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
[Crossref]

Pan, C. L.

J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
[Crossref]

Pedersen, L. H.

Peng, J. L.

Pfeiffer, H.

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

Pilevar, S.

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

Pollet, J.

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

Qu, H.

J. Li, H. Qu, and M. Skorobogatiy, “Detection of analyte refractive index and concentration using liquid-core photonic Bragg fibers,” Proc. SPIE 9702, 97020R (2016).
[Crossref]

J. Li, H. Qu, and M. Skorobogatiy, “Simultaneous monitoring the real and imaginary parts of the analyte refractive index using liquid-core photonic bandgap Bragg fibers,” Opt. Express 23(18), 22963–22976 (2015).
[Crossref] [PubMed]

H. Qu and M. Skorobogatiy, “Resonant bio- and chemical sensors using low-refractive-index-contrast liquid-core Bragg fibers,” Sens. Actuators B Chem. 161(1), 261–268 (2012).
[Crossref]

H. Qu, B. Ung, M. Roze, and M. Skorobogatiy, “All photonic bandgap fiber spectroscopic system for detection of refractive index changes in aqueous analytes,” Sens. Actuators B Chem. 161(1), 235–243 (2012).
[Crossref]

H. Qu and M. Skorobogatiy, “Liquid-core low-refractive-index-contrast Bragg fiber sensor,” Appl. Phys. Lett. 98(20), 201114 (2011).
[Crossref]

Riishede, J.

Rindorf, L.

Roze, M.

H. Qu, B. Ung, M. Roze, and M. Skorobogatiy, “All photonic bandgap fiber spectroscopic system for detection of refractive index changes in aqueous analytes,” Sens. Actuators B Chem. 161(1), 235–243 (2012).
[Crossref]

Sharma, A. K.

A. K. Sharma, R. Jha, and B. D. Gupta, “Fiber-optic sensors based on surface plasmon resonance: a comprehensive review,” IEEE Sens. J. 7(8), 1118–1129 (2007).
[Crossref]

Shiloach, M.

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

Shopova, S. I.

X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
[Crossref] [PubMed]

Sigalas, M.

Sirkis, J. S.

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

Skorobogata, O.

Skorobogatiy, M.

J. Li, H. Qu, and M. Skorobogatiy, “Detection of analyte refractive index and concentration using liquid-core photonic Bragg fibers,” Proc. SPIE 9702, 97020R (2016).
[Crossref]

J. Li, H. Qu, and M. Skorobogatiy, “Simultaneous monitoring the real and imaginary parts of the analyte refractive index using liquid-core photonic bandgap Bragg fibers,” Opt. Express 23(18), 22963–22976 (2015).
[Crossref] [PubMed]

H. Qu and M. Skorobogatiy, “Resonant bio- and chemical sensors using low-refractive-index-contrast liquid-core Bragg fibers,” Sens. Actuators B Chem. 161(1), 261–268 (2012).
[Crossref]

H. Qu, B. Ung, M. Roze, and M. Skorobogatiy, “All photonic bandgap fiber spectroscopic system for detection of refractive index changes in aqueous analytes,” Sens. Actuators B Chem. 161(1), 235–243 (2012).
[Crossref]

A. Mazhorova, A. Markov, A. Ng, R. Chinnappan, O. Skorobogata, M. Zourob, and M. Skorobogatiy, “Label-free bacteria detection using evanescent mode of a suspended core terahertz fiber,” Opt. Express 20(5), 5344–5355 (2012).
[Crossref] [PubMed]

H. Qu and M. Skorobogatiy, “Liquid-core low-refractive-index-contrast Bragg fiber sensor,” Appl. Phys. Lett. 98(20), 201114 (2011).
[Crossref]

A. Hassani and M. Skorobogatiy, “Photonic crystal fiber-based plasmonic sensors for the detection of biolayer thickness,” J. Opt. Soc. Am. B 26(8), 1550–1557 (2009).
[Crossref]

M. Skorobogatiy, “Efficient antiguiding of TE and TM polarizations in low-index core waveguides without the need for an omnidirectional reflector,” Opt. Lett. 30(22), 2991–2993 (2005).
[Crossref] [PubMed]

Smith, C.

Sun, C. K.

C. H. Lai, B. You, J. Y. Lu, T. A. Liu, J. L. Peng, C. K. Sun, and H. C. Chang, “Modal characteristics of antiresonant reflecting pipe waveguides for terahertz waveguiding,” Opt. Express 18(1), 309–322 (2010).
[Crossref] [PubMed]

J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
[Crossref]

Sun, Y.

X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
[Crossref] [PubMed]

Suter, J. D.

X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
[Crossref] [PubMed]

Tjin, S. C.

W. B. Ji, S. C. Tjin, B. Lin, and C. L. Ng, “Highly sensitive refractive index sensor based on adiabatically tapered microfiber long period gratings,” Sensors (Basel) 13(10), 14055–14063 (2013).
[Crossref] [PubMed]

Ung, B.

H. Qu, B. Ung, M. Roze, and M. Skorobogatiy, “All photonic bandgap fiber spectroscopic system for detection of refractive index changes in aqueous analytes,” Sens. Actuators B Chem. 161(1), 235–243 (2012).
[Crossref]

Uttamchandani, D.

A. Dudus, R. Blue, and D. Uttamchandani, “Comparative study of microfiber and side-polished optical fiber sensors for refractometry in microfluidics,” IEEE Sens. J. 13(5), 1594–1601 (2013).
[Crossref]

Wang, S. C.

T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013).
[Crossref] [PubMed]

Wang, X. D.

X. D. Wang and O. S. Wolfbeis, “Fiber-optic chemical sensors and biosensors (2008-2012),” Anal. Chem. 85(2), 487–508 (2013).
[Crossref] [PubMed]

West, J.

Wevers, M.

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

White, I. M.

X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
[Crossref] [PubMed]

Wolfbeis, O. S.

X. D. Wang and O. S. Wolfbeis, “Fiber-optic chemical sensors and biosensors (2008-2012),” Anal. Chem. 85(2), 487–508 (2013).
[Crossref] [PubMed]

Wu, C. C.

T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013).
[Crossref] [PubMed]

You, B.

Yu, C. P.

B. You, J. Y. Lu, J. H. Liou, C. P. Yu, H. Z. Chen, T. A. Liu, and J. L. Peng, “Subwavelength film sensing based on terahertz anti-resonant reflecting hollow waveguides,” Opt. Express 18(18), 19353–19360 (2010).
[Crossref] [PubMed]

J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
[Crossref]

Zhang, Z.

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

Zhu, H.

X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
[Crossref] [PubMed]

Zourob, M.

Anal. Bioanal. Chem. (1)

G. Gauglitz, “Direct optical sensors: principles and selected applications,” Anal. Bioanal. Chem. 381(1), 141–155 (2005).
[Crossref] [PubMed]

Anal. Chem. (3)

X. D. Wang and O. S. Wolfbeis, “Fiber-optic chemical sensors and biosensors (2008-2012),” Anal. Chem. 85(2), 487–508 (2013).
[Crossref] [PubMed]

M. P. DeLisa, Z. Zhang, M. Shiloach, S. Pilevar, C. C. Davis, J. S. Sirkis, and W. E. Bentley, “Evanescent wave long-period fiber bragg grating as an immobilized antibody biosensor,” Anal. Chem. 72(13), 2895–2900 (2000).
[Crossref] [PubMed]

T. C. Chang, C. C. Wu, S. C. Wang, L. K. Chau, and W. H. Hsieh, “Using a fiber optic particle plasmon resonance biosensor to determine kinetic constants of antigen-antibody binding reaction,” Anal. Chem. 85(1), 245–250 (2013).
[Crossref] [PubMed]

Anal. Chim. Acta (1)

X. Fan, I. M. White, S. I. Shopova, H. Zhu, J. D. Suter, and Y. Sun, “Sensitive optical biosensors for unlabeled targets: a review,” Anal. Chim. Acta 620(1-2), 8–26 (2008).
[Crossref] [PubMed]

Appl. Phys. Lett. (2)

H. Qu and M. Skorobogatiy, “Liquid-core low-refractive-index-contrast Bragg fiber sensor,” Appl. Phys. Lett. 98(20), 201114 (2011).
[Crossref]

J. Y. Lu, C. P. Yu, H. C. Chang, H. W. Chen, Y. T. Li, C. L. Pan, and C. K. Sun, “Terahertz air-core microstructure fiber,” Appl. Phys. Lett. 92(6), 064105 (2008).
[Crossref]

Biosens. Bioelectron. (1)

J. Pollet, F. Delport, K. P. Janssen, K. Jans, G. Maes, H. Pfeiffer, M. Wevers, and J. Lammertyn, “Fiber optic SPR biosensing of DNA hybridization and DNA-protein interactions,” Biosens. Bioelectron. 25(4), 864–869 (2009).
[Crossref] [PubMed]

IEEE Sens. J. (2)

A. K. Sharma, R. Jha, and B. D. Gupta, “Fiber-optic sensors based on surface plasmon resonance: a comprehensive review,” IEEE Sens. J. 7(8), 1118–1129 (2007).
[Crossref]

A. Dudus, R. Blue, and D. Uttamchandani, “Comparative study of microfiber and side-polished optical fiber sensors for refractometry in microfluidics,” IEEE Sens. J. 13(5), 1594–1601 (2013).
[Crossref]

J. Opt. Soc. Am. B (1)

Opt. Express (7)

A. Mazhorova, A. Markov, A. Ng, R. Chinnappan, O. Skorobogata, M. Zourob, and M. Skorobogatiy, “Label-free bacteria detection using evanescent mode of a suspended core terahertz fiber,” Opt. Express 20(5), 5344–5355 (2012).
[Crossref] [PubMed]

B. You, J. Y. Lu, J. H. Liou, C. P. Yu, H. Z. Chen, T. A. Liu, and J. L. Peng, “Subwavelength film sensing based on terahertz anti-resonant reflecting hollow waveguides,” Opt. Express 18(18), 19353–19360 (2010).
[Crossref] [PubMed]

C. H. Lai, B. You, J. Y. Lu, T. A. Liu, J. L. Peng, C. K. Sun, and H. C. Chang, “Modal characteristics of antiresonant reflecting pipe waveguides for terahertz waveguiding,” Opt. Express 18(1), 309–322 (2010).
[Crossref] [PubMed]

J. Li, H. Qu, and M. Skorobogatiy, “Simultaneous monitoring the real and imaginary parts of the analyte refractive index using liquid-core photonic bandgap Bragg fibers,” Opt. Express 23(18), 22963–22976 (2015).
[Crossref] [PubMed]

L. Rindorf, J. B. Jensen, M. Dufva, L. H. Pedersen, P. E. Høiby, and O. Bang, “Photonic crystal fiber long-period gratings for biochemical sensing,” Opt. Express 14(18), 8224–8231 (2006).
[Crossref] [PubMed]

A. Argyros, I. M. Bassett, M. A. Eijkelenborg, and M. C. J. Large, “Analysis of ring-structured Bragg fibres for single TE mode guidance,” Opt. Express 12(12), 2688–2698 (2004).
[Crossref] [PubMed]

J. West, C. Smith, N. Borrelli, D. Allan, and K. Koch, “Surface modes in air-core photonic band-gap fibers,” Opt. Express 12(8), 1485–1496 (2004).
[Crossref] [PubMed]

Opt. Lett. (4)

Proc. SPIE (1)

J. Li, H. Qu, and M. Skorobogatiy, “Detection of analyte refractive index and concentration using liquid-core photonic Bragg fibers,” Proc. SPIE 9702, 97020R (2016).
[Crossref]

Sens. Actuators B Chem. (3)

A. Iadicicco, D. Paladino, S. Campopiano, W. J. Bock, A. Cutolo, and A. Cusano, “Evanescent wave sensor based on permanently bent single mode optical fiber,” Sens. Actuators B Chem. 155(2), 903–908 (2011).
[Crossref]

H. Qu and M. Skorobogatiy, “Resonant bio- and chemical sensors using low-refractive-index-contrast liquid-core Bragg fibers,” Sens. Actuators B Chem. 161(1), 261–268 (2012).
[Crossref]

H. Qu, B. Ung, M. Roze, and M. Skorobogatiy, “All photonic bandgap fiber spectroscopic system for detection of refractive index changes in aqueous analytes,” Sens. Actuators B Chem. 161(1), 235–243 (2012).
[Crossref]

Sensors (Basel) (1)

W. B. Ji, S. C. Tjin, B. Lin, and C. L. Ng, “Highly sensitive refractive index sensor based on adiabatically tapered microfiber long period gratings,” Sensors (Basel) 13(10), 14055–14063 (2013).
[Crossref] [PubMed]

Other (3)

M. Zourob, S. Elwary, and A. Turner, Principles of Bacterial Detection: Biosensors, Recognition Receptors and Microsystems (Springer, 2008).

M. Skorobogatiy, Nanostructured and Subwavelength Waveguide: Fundamentals and Applications (Wiley, 2012).

M. Skorobogatiy and J. Yang, Fundamentals of Photonic Crystal Guiding (Cambridge University Press, 2009).

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Figures (11)

Fig. 1
Fig. 1 Schematic of the Bragg fiber (a) and the corresponding planar Bragg waveguide (b) with the same material composition and layer thickness. The Bragg reflectors feature alternating polystyrene (PS) /poly-methacrylate (PMMA) layers. The green region represents the analyte layer.
Fig. 2
Fig. 2 Simulated modal dispersion relations of TM1 (a) and TE1 (b) modes in the water-filled Bragg fiber featuring an analyte layer with thickness ranging from 0nm to 200nm. Corresponding radiation losses of the TM1 (c) and TE1 (d) modes. Dependence of the wavelength of avoided crossing of the TM1 (e) and TE1 (f) modes (obtained from (c) and (d)) as a function of the analyte layer thickness. Dependence of the wavelength of the lowest radiation loss of the TM1 (g) and TE1 (h) modes (obtained from (c) and (d)) as a function of the analyte layer thickness.
Fig. 3
Fig. 3 Schematic of the modal dispersion relations of the core guided (blue dotted line) and analyte layer bound modes (purple dotted line), as well as dispersion relations of the hybridized modes in a Bragg fiber with an analyte layer (green solid lines) near the point of phase matching (anticrossing wavelength).
Fig. 4
Fig. 4 Simulated surface sensitivity of 10 even symmetric TE (a) and TM (b) modes (closest to the light line) when the fiber core size D is reduced from 600μm to 20μm. The PVB layer thickness is kept constant and equal to 100nm in all the simulations.
Fig. 5
Fig. 5 (a) Cross section of the hollow-core PMMA/PS Bragg fiber with a PVB layer under optical microscope. (b) Cross section of the Bragg reflector taken by a scanning electron microscope (SEM), featuring the alternating polystyrene (PS) and poly-methacrylate (PMMA) layers with individual thickness of ~210nm, as well as a ~500nm-thick PVB layer.
Fig. 6
Fig. 6 Experimental setup for the optical characterization of the surface sensitivity of the squeezed Bragg fiber sensor. The 8cm-long Bragg fiber with a thin PVB coated on the inner surface is integrated into the set-up using two opto-fluidic blocks. A broadband supercontinuum beam is launched into one end of the liquid-core Bragg fiber using an objective, while the output spectrum of the fiber sensor is registered by a monochromator. The metallic rod (diameter: 2.54cm) is fixed on a micro-positioning stage and is used to squeeze a small section of the fiber in increments of 50μm. Inset: ~8cm Bragg fiber used in this experiment (top), squeezing sketch of our setup (middle), as well as the squeezed section of the Bragg fiber (bottom).
Fig. 7
Fig. 7 Experimental transmission spectra of the Bragg fiber during the dissolution process of a thin PVB layer (thickness~500nm) coated on the inner surface of the fiber core. A blue shift is observed with dissolution of the PVB film. After the PVB layer is dissolved, the fiber is purged with fresh ethyl alcohol, and a red shift of ~5nm is observed, which is attributed to the fact that the RI of ethyl alcohol is somewhat lower than that of the PVB-alcohol mixture.
Fig. 8
Fig. 8 (a) Experimental transmission spectra of the ~8cm-long Bragg fiber filled with distilled water (a), (c), and 16%wt NaCl in water (b), (d), before and after the PVB film dissolution. The ~500nm PVB film induces a red shift of 26nm and 23nm in the transmission peak, respectively, when compared to transmission of the same fiber without a PVB layer.
Fig. 9
Fig. 9 Experimental transmission spectra of the ~8cm long Bragg fiber filled with distilled water before PVB layer dissolution, when a section of fiber is squeezed by displacements ranging from 0μm to 500μm using a 2.54cm diameter cylinder. The fiber has a PVB layer with thickness of 500nm.
Fig. 10
Fig. 10 Experimental transmission spectra of the ~8cm long Bragg fiber filled with distilled water and 16%wt NaCl solution in water before and after the PVB film dissolution. The fiber is squeezed by the depths of 250μm and 500μm using a 2.54cm diameter cylinder. In the case of a 250μm-squeezed fiber (D~350μm), a ~500nm PVB film induces a red shift of 30nm (filled with pure water), and 26nm (filled with 16%wt NaCl solution). In the case of a 500μm-squeezed fiber (D~100μm), the 500nm PVB film induces a red shift of 35nm (filled with pure water), and 29nm (filled with 16%wt NaCl solution).
Fig. 11
Fig. 11 Surface sensitivities of the squeezed hollow core Bragg fiber sensor filled with distilled water and 16%wt NaCl solution for various degrees of squeezing. The circular fiber core has a diameter of ~600μm. Inset: Schematic of the squeezed Bragg fiber with core size of D.

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