Abstract

We propose, numerically analyze and experimentally demonstrate a novel refractive index sensor specialized for low index sensing. The device is based on a directional coupler architecture implemented in a single microstructured polymer optical fiber incorporating two waveguides within it: a single-mode core and a satellite waveguide consisting of a hollow high-index ring. This hollow channel is filled with fluid and the refractive index of the fluid is detected through changes to the wavelength at which resonant coupling occurs between the two waveguides. The sensor design was optimized for both higher sensitivity and lower detection limit, with simulations and experiments demonstrating a sensitivity exceeding 1.4 × 103 nm per refractive index unit. Simulations indicate a detection limit of ~2 × 10−6 refractive index units is achievable. We also numerically investigate the performance for refractive index changes localized at the surface of the holes, a case of particular importance for biosensing.

© 2014 Optical Society of America

1. Introduction

Fiber-optic refractive index (RI) sensors have been widely studied in recent years for diverse applications in environmental monitoring, metrology, chemical and bio-medical sensing [1]. Using optical fibers rather than bulk optics or planar waveguide approaches offers the potential advantage of cost-effective mass-fabrication and ease of alignment. Various types of fiber-based RI sensor have been demonstrated based on surface plasmon resonance (SPR) [24], long period gratings (LPGs) [5, 6], photonic crystal fibers (PCFs) [79], fiber Fabry-Perot cavities [10, 11] and so on. PCF-based RI sensors, relying on the directional coupling between a single core mode and the modes of a fluid channel have proven to have particularly low detection limits, and because they do not rely on longitudinal patterning, could in principle be mass-produced by the fiber drawing technique [7, 1214]. However, our previous demonstrations suffered several disadvantages for practical applications: First, the sensor required a labor-intensive selective filling of liquid analyte into one of many micron sized air-holes for each measurement. Second, it relied on using analytes with refractive index higher than that of silica. In bio-sensing, analytes with bio-molecules are mostly given as water-based solutions whose indices are around 1.33, so that RI sensors operating in the low-index region are highly desirable for bio-medical applications [15]. More recent types of fiber directional coupler sensors have been reported which avoid the need of selective filling, including fibers with a single analyte channel [12], but the latter cannot be applied to low refractive index analytes either.

A number of fiber-based RI sensors specifically designed for low-index measurement were proposed recently, for example based on LPGs [16, 17], etched fiber-loops [18], or twin-core PCF [9, 14]. However, most of these devices require complicated post-processing such as irradiation by CO2 laser or fs-laser pulses for inscribing the LPGs on the sensor fibers [16, 17], partial etching of unclad-fiber [18], and selective filling [14]. A promising approach developed by Frosz et al. working for low refractive indices and requiring no post-processing [19], exploits the sensitivity of phase matching of four-wave mixing to the refractive index in the holes of a PCF, with detection limits better than 6 × 10−6 deemed possible, although the nonlinear optofluidic setup might prove difficult to use in practice. To our knowledge a linear optics RI sensor based on fiber without requiring post-processing, working for low refractive indices, with detection limits that could compete with surface plasmon resonance sensors is still lacking. The dual core architecture described in [20] appears to fulfil all these conditions, but to our knowledge has not yet been demonstrated in practice.

In this paper, we propose and demonstrate a novel RI sensor not requiring any complicated post-processing and suitable for low-index sensing. Our device is based on a directional coupler implemented in a single micro-structured fiber as in [7, 13], but its structure is different in that the sensor fiber includes only a single micrometric hollow fluid channel instead of a holey-cladding. In that, the geometry is similar to the single fluid channel directional coupler described in [12], but with an additional thin high-index ring surrounding the fluid channel [21] and a composite core engineered to provide suitable phase matching while remaining single-moded with a relatively large diameter for ease of coupling [22]. The dispersion and cutoffs of the modes of the high index ring depend on the refractive index of the fluid filling them, and previous studies have shown that a thin, high index ring provides the best sensitivity [21]. For the high index contrast in the channel, we consider a polymer-based fiber [23] rather than a silica-based one, since the high index contrast is challenging to achieve using glass fabrication techniques relying on doping, but can be obtained with polymer drawing techniques. The polymer-based fiber also provides several additional advantages: First, the polymers can readily capture organic compounds, enabling the sensor to be used for bio-sensing applications. Second, the polymer fiber is mechanically robust, so that it does not shatter nor produce harmful shards. In our device, three kinds of polymers are considered: a poly-(methyl methacrylate) (PMMA) background, a composite core consisting of Zeonex 480R and PMMA [22] and polycarbonate (PC) for the high-index ring.

2. Fiber-based directional coupler sensor for low-index bio-sensing

2.1 Geometry and principle of operation

Figure 1(a) and 1(b) illustrate the studied geometry. The polymer fiber includes an all-solid composite core and a hollow high-index ring satellite waveguide which may be filled with a liquid, low-index analyte. The composite core consists of an array of higher refractive index polymer rods (Zeonex [24]) in the PMMA background, designed for single-mode operation, as detailed in Section 3.1. Light propagating in the core can couple to the satellite waveguide, with full power transfer possible when a mode of the satellite waveguide is phase matched to a core mode. For a given satellite waveguide mode, phase matching only occurs at a resonant wavelength (λr) where nco(λr) = nsat(λr), where nco and nsat denote the effective index of the core and satellite modes, respectively [see Fig. 1(c)] [7,13,25]. At the resonant wavelength, power is transferred periodically along the coupler between the core and satellite mode, with complete power transfer occurring at odd multiples of the coupling length Lc. If the length of fluid Lfluid in the satellite channel is close to a multiple of Lc, a dip in the core’s transmission spectrum is observed at the resonant wavelength. Since the optical field of the satellite mode has a strong overlap with the analyte, nsat depends on the analyte’s refractive index na, and so does the resonant wavelength. By measuring the transmission spectrum through the fiber the refractive index can thus be measured. The difficulty in the design is to achieve suitable coupling lengths, and suitable dispersion curves for the sensor to work with aqueous analytes, in a geometry that can realistically be fabricated and operated.

 figure: Fig. 1

Fig. 1 (a) Schematic diagram of the directional coupler geometry within a polymer fiber. (b) Details of the cross section; (c) Schematic of effective index curves of core and satellite modes and resonance wavelength.

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2.2 Refractive index sensitivity and detection limit

A detailed analysis of fiber directional couplers used as a refractive index sensor and their performance can be found in [13]. Here we present only the essential results derived in [13] necessary for the assessment of the device under investigation.

The performance of the sensor can be described in terms of the sensitivity S (shift in resonance wavelength per change in refractive index) or detection limit δn (minimum detectable change in refractive index). The detection limit is ultimately the only quantity of concern in practice, and while it is inversely proportional to the sensitivity it also depends on the spectral width of the dip to be detected and the signal to noise ratio [26]. The width of the dip and sensitivity are not completely independent quantities, and improving one by design often is compensated by deterioration in the other. If Lfluid is not close to Lc, the power transfer between core and satellite waveguide is limited, resulting in a finite depth of the transmission dip that is typically more limiting than the signal to noise ratio. Ultimately, the detection limit δn can be approximated by [13]

δnαλrLcfsatTmin(SNR)0.25,
where α is a number close to 0.7 depending on the group index of the satellite mode and the refractive index of the analyte, Tmin is the relative depth of the transmission dip (ratio of transmission at the minimum of the dip to maximum transmission near the dip) and is largely determined by how close Lfluid is to Lc, fsat is an overlap integral quantifying the proportion of the satellite field intensity overlapping with the analyte, and SNR is the signal to noise ratio in linear units [13]. Explicitly,
fsat=analyteεsat(x,y)|Esat(x,y)|2dxdycrosssecsionεsat(x,y)|Esat(x,y)|2dxdy,
where εsat is the permittivity distribution of the satellite waveguide, and Esat the modal field. As Tmin and SNR are largely dependent on the quality of the experimental setup, the optimization of a design is limited to the optimization of Lc and fsat. The coupling length Lc is directly related to the overlap integral between the core mode and the satellite mode, and thus strongly dependent on the distance between the composite core and analyte channel. Clearly the larger Lc the better detection limit, but in practice filling sections much longer than a few centimeters of micrometric capillaries can become impractical due to capillary forces. The overlap integral fsat depends solely on the geometry of the analyte channel, which in the present case is defined by the radius of the hole, rin, and the high index ring’s thickness t and refractive index. As previous studies have shown high refractive index rings are the most promising, we use polycarbonate, which has a refractive index ~0.08 higher than PMMA, leaving only t and rin to be optimized.

3. Modeling of the sensor fiber and its optimization for low-index bio-sensing

3.1 Fiber design

For well-distinguishable resonances, a single-mode fiber core is highly desirable. However, polymer fibers typically have relatively large index contrasts, for which single-mode cores would be very small and difficult to couple into through butt coupling from conventional fibers. We solve this issue by employing a recently proposed composite core composed of an array of 19 Zeonex rods (n ~1.52) with hexagonal symmetry in a PMMA (n ~1.486) background [22]. Figure 2(a) illustrates the schematic design of the composite core, where the diameter of each rod (d) and the lattice period (or pitch, Λ) are 345 nm and 860 nm, respectively. Figure 2(b) shows the intensity profile of the lowest order core mode simulated using the multipole method at the wavelength of 800 nm [27], at which optical sources are readily available and water has good transparency. The array of rods effectively acts as a single-mode core with low index contrast (~0.1%), resulting in a larger core mode (field diameter of 5.0 μm) [22] as compared to the conventional step-index cores using either Zeonex or PC which would require a core diameter of 1.9 μm and 1.2 μm respectively to remain single-mode as shown in Fig. 2(c) and 2(d).

 figure: Fig. 2

Fig. 2 (a) Schematic design of the composite core. The intensity profiles of the lowest order modes in (b) composite and circular single-mode cores using (c) Zeonex and (d) PC. All cases shown support two degenerate modes of orthogonal polarization: LP01x and LP01y (or HE11x and HE11y).

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On the contrary, a low index contrast at the satellite waveguide is not desirable because it increases the spread of modal field into the background material and thus reduces fsat. We consider here the analyte channel defined by a ring of PC (n ~1.57) in a PMMA background (n ~1.486), and in this case, the index contrast between the ring and background reaches 6%.

Figure 3(a) illustrates the schematic design of the satellite waveguide, where t and rin are the thickness and inner radius of the channel, respectively. In order to obtain well-defined resonances, the satellite waveguide should also be as few-moded as possible, which requires t and rin to both be small. The number of ring modes increases rapidly with t and rin and it is thus desirable to keep both values as small as possible or practicable. The hole should be also large enough to be readily filled with analyte, ideally simply through capillary forces. This is achievable by employing a hole with a radius of 2 μm that is surrounded by a 300 nm thick PC ring: The calculated minimum t to support modes around 800 nm is slightly smaller than 300 nm, but for the sake of convenience in fiber fabrication, we set the thickness to be 300 nm. With these dimensions, the satellite waveguide supports only four guided modes around 800 nm wavelength: TM, TE, and two HE degenerate mode pairs, and their intensity profiles simulated for a low-index analyte (na = 1.333) are shown in Figs. 3(b)-3(e). The modal overlap in the analyte hole (fsat) is also affected by the hole-size, which gives the lower limit of rin as will be discussed in Section 3.3.

 figure: Fig. 3

Fig. 3 (a) Schematic design of the satellite waveguide (rin = 2 μm, t = 300 nm). The intensity profiles of four guided modes in the satellite waveguide when the hole coated with a PC ring is filled with low-index analyte (na = 1.333): (b) TM, (c) TE, (d) HE-1 (e) HE-2 modes.

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3.2 Directional coupling in the designed fiber

Figure 4(a) shows the calculated effective index curves for the interacting modes discussed in the previous section. The resonant coupling between the core and four satellite modes occurs at each crossing satisfying the phase matching condition, nco = nsat, as discussed in section 2.1.

 figure: Fig. 4

Fig. 4 (a) Calculated effective index for the core and four satellite modes showing the resonant wavelength, and (b) the coupling length for HE11x-TM mode combination plotted as a function of the center-to-center distance (h) between the core and satellite waveguides. The distance, h, is defined in the inset of Fig. 4(b).

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Due to the evanescent decay of fields outside the waveguide, the coupling length is an almost exponential function of the distance between two waveguides. Figure 4(b) shows an example of Lc for the HE11x-TM mode combination plotted as a function of the center-to-center distance (h) between the core and satellite waveguides. The calculated value of Lc for h = 10 μm is 1.65 cm, which is suitable to realize cost-effective and disposable devices. The coupling lengths for two degenerate core and four satellite mode combinations at each resonance are summarized in Table 1.For h = 10 μm, the calculated values of Lc for all combinations are less than 2 cm, allowing practical applications of the proposed sensor, although if filling ~10 cm is not a problem longer coupling lengths obtained for slightly larger h values would improve the detection limit accordingly.

Tables Icon

Table 1. Coupling lengths for each core-satellite mode combination at the corresponding resonant wavelengths, for h = 10 μm. Only one polarization for each degenerate pair shown.

3.3 Low-index sensing and its optimization

Sensor performance is closely related to the overlap integral fsat, which for given materials and thickness depends on hole size and wavelength. The fact fsat is a function of wavelength can be understood in that for λ → 0 fields are confined to the high index ring, so that fsat → 0, while for λ >> rin the modal fields expand into the cladding and fsat again tends to zero. In between these asymptotes there must thus lie a wavelength of maximum fsat optimizing sensing. Figure 5(a) shows the calculated fsat as a function of wavelength for the TM and TE satellite modes. The corresponding curves for the two HE modes show similar trends and are thus not included. The results show that fsat for the TM mode at the resonance for the HE11x-TM mode pair (λr = 746 nm) is larger than for the TE mode at the corresponding resonance (λr = 861.8 nm). The overlap fsat for TM mode reaches 3.28% at λ = 746 nm, and the corresponding sensitivity and detection limit are calculated to be 1.2 × 103 nm/RIU and ~3.6 × 10−6 RIU, respectively. Here, the heuristic values of SNR ( = 36dB) and Tmin ( = −30dB) taken from [13] are used for the calculation. Figure 5(b) shows fsat(λ) for the TM mode for several different sizes of the analyte hole. Curves become overlapping for large values of rin, showing the value of fsat converges with increasing radius – this can be understood in that for large radii curvature becomes irrelevant and modes are locally essentially those of a thin slab [28]. The hole-radius, rin, should thus be large enough for fsat to have converged to its maximum. We thus chose an inner radius of 2 μm to maximize fsat while maintaining small sample volume. Note that alternatively larger diameters could be chosen without impacting fsat, and potentially making filling easier, but this comes at a cost of a higher density of azimuthal modes, and thus a more complex transmission spectrum. The radius of 2 μm simultaneously satisfy the requirements for maximizing fsat, maintaining fewest 4 satellite modes, and ease in infiltration of analyte through capillary forces.

 figure: Fig. 5

Fig. 5 (a) Calculated modal intensity overlap in the analyte hole (fsat) for TM and TE modes, and (b) fsat curves for TM mode plotted for different sizes of the analyte hole.

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As shown in Fig. 5(a), the maximum value of fsat for the TM mode is calculated to be 4.69% at λ = 600 nm, leading to a sensitivity of 1.4 × 103 nm/RIU and detection limit of ~2 × 10−6 if operating at that wavelength and with the same coupling length of 1.65 cm (note this would require redesigning the core to obtain the desired resonant wavelength, and select h accordingly). The detection limits for fiber-based low-index sensing reported in recent publications are listed in Table 2.The predicted detection limit of our device is comparable to the best results, which means that the proposed device has a very good potential for low-index sensing with a relatively simple fabrication and measurement setup.

Tables Icon

Table 2. Detection limits for fiber-based low-index sensing reported in recent publications.

3.4 Performance for surface refractive index changes and biosensing

One of the main motivations behind integrated high sensitivity refractive index sensors is label free detection of biomolecules, which is achieved by attaching selective molecules such as antigens or antibodies to the inside of the microfluidic channel, and detecting changes in local refractive index due to binding events with the target molecule. The refractive index change in that case is localized to a thin layer inside the channel [20,29], and the detection limit for bulk refractive index changes is no longer relevant as such. For the present geometry, fields decay evanescently away from the high index layer so that the sensitivity at its surface will contribute strongly to the overall response.

To estimate the potential of our sensor for biosensing, we consider an idealized model of detection of α-streptavidin, using a streptavidin antigen layer immobilized on the inner surface of channel. A streptavidin molecule has a typical size of Vmol = 5.5 × 4.5 × 4.5 nm, and its molecular mass (mmol) is 9.96 × 10−23 kg [30]. We consider the refractive index of protein and the antigen to be 1.45, and treat the layer as a homogenous layer with an average thickness tb = σVmol/mmol, where sigma is the surface concentration of the protein. The sensitivity can now be defined as the shift of resonant wavelength per change in effective protein layer thickness (S' = r/dtb) and the detection limit as the minimum detectable change in tb or corresponding concentration assuming all proteins get captured. We calculate the sensitivity using full multipole simulations of the optimized fiber from the previous section but with an antigen layer 20 nm thick deposited in the hole. We find S' = 1.23 nm/nm, which in ideal conditions (Lfluid = Lc, SNR = 36dB, Tmin = −30dB) corresponds to a minimal detectable change of tb of only 3.4 pm corresponding to a concentration of 3.1 μg/mL. We emphasize this is an ideal scenario. Achieving such low detection limits would require detecting a wavelength shift in transmission dip of only 3 pm which can only be achieved with great difficulty in practice. The previously reported detection limits for fiber-based protein measurement are listed in Table 3.The predicted detection limit of our device records one of the best results, which means that the proposed device also has a good potential for bio-sensing, arguably with a relatively simple device implementation. Thanks to the small size of channel, the proposed technique allows detecting bio-molecules using small analyte volumes (down to 0.2 nL in our case) and thus the fabrication of miniaturized portable devices.

Tables Icon

Table 3. Detection limits for fiber-based protein measurement reported in recent publications.

4. Fabrication and characterization

The fiber was fabricated using a multi-stage polymer fiber drawing technique. The composite core was fabricated as in [22], and drawn to an outer diameter (OD) of 362 μm. Separately, a cane was fabricated by first drawing a PMMA preform (OD 68 mm) containing one central hole (diameter) 2.5 mm and six satellite holes (1.6 mm, 3.4 mm center to center distance from central hole) down to an OD of 10 mm. One of the satellite holes will be the analyte channel, while the additional five holes are introduced at equal distance from the core in a hexagonal arrangement to increase symmetry and minimize deformations during the initial drawing stages. A PC capillary (OD 230 μm, wall thickness 15 μm) is inserted in one of the six satellite holes, becoming the analyte channel surrounded by a high index PC ring during the fiber drawing process, as designed in section 3.1. Five PMMA rods (OD 230 μm) are also inserted in the other five holes to retain constant cross-sectional strain on the fiber so that the inner micro-structure of the sensor fiber is maintained to be uniform during the whole drawing process. The 362 μm composite core is inserted in the central hole of this cane. At this cane stage, the cross section is very close to a 57.5 times larger version of that of the optimized design presented in Section 3.3. The cane was then drawn in a two stage process down to a range of ODs between 150 μm and 300 μm by varying the drawing speeds.

A range of final dimensions was sought as the coupling wavelength and coupling length derived in Section 3.3 are exquisitely sensitive to details of the geometry. Selection of an appropriate stretch of fiber was done first through visual inspection of the cross section (to ensure appropriate dimensions and that the PC hole had not collapsed). The most promising stretches were then tested for transmission dips without filling the fiber. Finally, the most suitable fiber, shown in Fig. 6(a), was characterized in a transmission measurement with aqueous sucrose solutions of various concentrations as an analyte. The section of fiber used in measurements had an outer diameter of 198 μm, h = 10.0 μm [both measured using a scanning electron microscope (SEM)], with optical micrographs showing the composite core to have Λ = 1.0 ± 0.2μm. Because of the weak contrast between Zeonex and PMMA, the diameter of the Zeonex rods can only be estimated from optical micrograph and the initial d/Λ ratio of d = 360 ± 70nm [22]. Using the cutback technique we measured propagation losses in the core to be 0.28dB/cm at 550nm and 0.34dB/cm at 750nm, which is about 4 times higher than the same core in [22] – likely due to darkening of the Zeonex with time that occurred in the two years since the composite core had been first drawn. Figure 6(b) shows the spectral dependence of the transmission of the empty fiber over the region of interest, normalized to its maximum to remove input and output coupling losses.

 figure: Fig. 6

Fig. 6 (a) Microscope image of the core and analyte channel of the fiber; (b) transmission spectra through the empty fiber (blue), and after infiltration for various concentration of sucrose, as indicated – spectra are offset for readability; (c) Measured wavelength of the dips indicated by red arrows in (b) as a function of refractive index. The error bars reflect the accuracy of the refractometer ( ± 0.0003) and of the OSA ( ± 0.1 nm).

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Using capillary forces to fill the minute fiber hole was not sufficient because of the low wettability of PC – this would need to be improved for any practical applications. The analyte was thus injected in the satellite core by a high pressure chromatography pump (Alltech Model 426) through an aluminum pressure chamber with an optical window. The satellite core is infiltrated over a length of 9 cm (corresponding to an analyte volume of approximately 1.1 nL). This infiltration length, which is several times greater than the expected coupling length, was the shortest we could achieve with the current design of our pressure chamber. Over the wavelength range of interest, the core is single-moded and the absorbance of water and sucrose solution is negligible [34].

Light from a super-continuum (SC) source (Crystal Fiber: Pavg < 100 mW, spectral range = 500 – 1750 nm) is launched into the composite core through a short pass filter (λc = 900 nm) to eliminate the higher powered long wavelengths of the spectrum which could damage the fiber. The output from the other fiber-end is monitored with an optical spectrum analyzer (OSA) (see Fig. 5 showing the schematic of experimental setup in [13]). The output light from the satellite core is spatially blocked to keep only light from the composite core, and an output polarizer aligned with the direction of the composite core and satellite core [i.e. horizontally on Fig. 6(a), corresponding to the HE1x mode] is used to select the relevant combination of the interacting modes. From Table 1 if the sample had been perfectly symmetric, the only ring modes that could be coupled to would be TM and HE-2x. Because of the slight offset of the satellite mode with respect to the x axis of the core [Fig. 6(a)], coupling is also possible to the HE-2y mode, with slightly different resonant wavelength due to the asymmetry. The resulting spectra are measured with water (refractive index is 1.3334), 4% (1.3381), 5% (1.3392), 6% (1.3414) and 7% (1.3430) sucrose solution, respectively. We use a refractometer (Atago PAL-RI, Accuracy ± 0.0003) to verify the refractive index of the solution. The aluminium reservoir is rinsed thoroughly and dried each time the solution is changed, and the next solution flushed through the fiber for several minutes at each step. Figure 6(b) shows the spectra for the various concentrations, with vertical shifts between each curve proportional to the refractive index differences to guide the eye.

A shortcoming of this experiment is that Lfluid is fixed to the total length of the fiber, and thus an unknown multiple of the coupling length of the modes, given the sensitivity of coupling length to the exact structure and h in particular. The spectrum does nonetheless show four distinct transmission dips, of which two are over 7dB in depth. The three dips at the longest wavelengths correspond to the expected TM and HE-2x,y modes, while the shortest wavelength dip can be attributed to an HE-3 mode that was ignored in simulations because its resonant wavelength was below the studied wavelength range for water infiltration.

While we were unable to get an accurate direct measurement of the PC ring thickness using electron microscopy because of the vanishing contrast between PC and PMMA, the resonant wavelengths are in agreement with simulations for a ring thickness between 250 and 300 nm; resonances would exist up to 1μm wavelength if the thickness was 350 nm, but no resonances were observed above 800 nm.

Figure 6(b) shows the wavelength of the two deepest resonances as a function of the analyte’s refractive index. The slope of each curve yields a sensitivity for each dip of S = 1.43 × 103 nm/RIU (blue) and S = 1.66 × 103 nm/RIU (red), in relatively good agreement with simulations, given the manufactured fiber’s cross section does not match the design perfectly. The high sensitivity is thus confirmed, and lies among the highest demonstrated in a fiber based refractive index sensor for low index sensing to date. However the limited depth of the dips measured because of the fixed length of the fiber sample does not allow us to reach the best possible detection limits with this geometry. In our experiment the detection limit is limited by the resolution of the OSA to 0.1nm, which once divided by the sensitivity 1.66 × 103 corresponds to 6 × 10−5 RIU. This is simply the resolution with which we can determine the minimum in the spectrum experimentally on the OSA. The formalism of Eq. (1) strictly only applies for single, well isolated dips – in particular the width of dips becomes ill-defined in the presence of two close dips. The most isolated dip in our measurement (the TM mode for 4% sucrose solution), has a full width 3dB above the minimum of 10.6 nm. Using our measured Tmin = −7 dB, SNR = 27 dB and the simulated values of fsat, λr and Lc from Section 3 yields a more conservative value of 1.0 × 10−4 RIU.

5. Conclusion and discussion

In conclusion, we presented the design, fabrication and characterization of a fiber-based refractive index sensor working for low index analytes, based on a directional coupler implemented in a polymer composite fiber including only a single analyte channel coated with a high-index polymer. We measured a sensitivity of 1.66 × 103 nm/RIU and the detection limit for low-index sensing is predicted to reach ~2 × 10−6. We showed the geometry could in principle be used for bio-sensing, with a detection limit of order 3 μg/mL. The results are comparable to the best fiber sensing performance reported in recent work. However, unlike most fiber sensing schemes requiring gratings, interferometers or selectively filled PCFs, the sensor we presented here consist of a simple length of fiber, not requiring any post-processing, and is thus easily mass-produced.

Nonetheless, several issues will need to be resolved for this sensor to become really practical: ideally the holes could be filled using capillary forces, which will require a surface that is more hydrophilic than polycarbonate. This could be achieved through surface treatment, possibly at the preform stage given the relatively low drawing temperatures. Furthermore, PMMA can absorb up to 2% by weight of water, with an index change up to 0.002. The diffusion of water through the end face is too slow to affect measurement, but diffusion from the capillary to the region between capillary and core will be relevant even over timescales of a minute or less, adding a time dependent drift in the dip’s depth and wavelength. The PC ring is expected to slow the diffusion into PMMA from the capillary somewhat, and we have not observed such a drift in our limited experiments, but it is clear that this effect will need to be taken into account to achieve the best possible performance.

Note that for biosensing, the inside of the hole would need coating with bioselective molecules, which requires flushing a series of solutions through the hole. Coating with a hydrophilic layer could then be done in the same process. Finally considerable improvement could be achieved by using the fiber in a reflection setup by adding a reflective layer to one fiber end facet [13] – this way light is injected and measured from the same fiber end, and the reflection spectrum can be monitored as the analyte fills the hole, which guarantees measurements can be obtained for Lfluid = Lc. An all-fiber setup as was used for example in [13] will however be difficult as there are currently no fiber circulators available for the wavelength range of interest.

Acknowledgment

This research was supported under the Australian Research Council’s (ARC) Discovery Project scheme (DP0881528). B. T. K., and A. A. acknowledge support from an ARC Future Fellowship, and an ARC Australian Research Fellowship, respectively. This work was performed in part at the OptoFab node of the Australian National Fabrication Facility, a company established under the National Collaborative Research Infrastructure Strategy to provide nanofabrication and micro-fabrication facilities for Australian researchers. Xiaoqi Liu acknowledges support from the China Scholarship Council.

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12. H. W. Lee, M. A. Schmidt, P. Uebel, H. Tyagi, N. Y. Joly, M. Scharrer, and P. St. J. Russell, “Optofluidic refractive-index sensor in step-index fiber with parallel hollow micro-channel,” Opt. Express 19(9), 8200–8207 (2011). [CrossRef]   [PubMed]  

13. D. K. C. Wu, K. J. Lee, V. Pureur, and B. T. Kuhlmey, “Performance of Refractive Index Sensors Based On Directional Couplers in Photonic Crystal Fibers,” J. Lightwave Technol. 31(22), 3500–3510 (2013). [CrossRef]  

14. B. Sun, M.-Y. Chen, Y.-K. Zhang, J.-C. Yang, J.-Q. Yao, and H.-X. Cui, “Microstructured-core photonic-crystal fiber for ultra-sensitive refractive index sensing,” Opt. Express 19(5), 4091–4100 (2011). [CrossRef]   [PubMed]  

15. 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]  

16. L. Rindorf and O. Bang, “Highly sensitive refractometer with a photonic-crystal-fiber long-period grating,” Opt. Lett. 33(6), 563–565 (2008). [CrossRef]   [PubMed]  

17. M. Yang, D. N. Wang, Y. Wang, and C. Liao, “Long period fiber grating formed by periodically structured microholes in all-solid photonic bandgap fiber,” Opt. Express 18(3), 2183–2189 (2010). [CrossRef]   [PubMed]  

18. C. Wang and C. Herath, “High-sensitivity fiber-loop ringdown evanescent-field index sensors using single-mode fiber,” Opt. Lett. 35(10), 1629–1631 (2010). [CrossRef]   [PubMed]  

19. M. H. Frosz, A. Stefani, and O. Bang, “Highly sensitive and simple method for refractive index sensing of liquids in microstructured optical fibers using four-wave mixing,” Opt. Express 19(11), 10471–10484 (2011). [CrossRef]   [PubMed]  

20. C. Markos, W. Yuan, K. Vlachos, G. E. Town, and O. Bang, “Label-free biosensing with high sensitivity in dual-core microstructured polymer optical fibers,” Opt. Express 19(8), 7790–7798 (2011). [CrossRef]   [PubMed]  

21. B. T. Kuhlmey, S. Coen, and S. Mahmoodian, “Coated photonic bandgap fibres for low-index sensing applications: cutoff analysis,” Opt. Express 17(18), 16306–16321 (2009). [CrossRef]   [PubMed]  

22. S. G. Leon-Saval, R. Lwin, and A. Argyros, “Multicore composite single-mode polymer fiber,” Opt. Express 20(1), 141–148 (2012). [CrossRef]   [PubMed]  

23. A. Argyros, “Microstructured polymer optical fibers,” J. Lightwave Technol. 27(11), 1571–1579 (2009). [CrossRef]  

24. http://www.zeonex.com/datasheets.asp

25. B. T. Kuhlmey, B. J. Eggleton, and D. K. C. Wu, “Fluid-filled solid-core photonic bandgap fibers,” J. Lightwave Technol. 27(11), 1617–1630 (2009). [CrossRef]  

26. I. M. White and X. D. Fan, “On the performance quantification of resonant refractive index sensors,” Opt. Express 16(2), 1020–1028 (2008). [CrossRef]   [PubMed]  

27. B. T. Kuhlmey, T. P. White, G. Renversez, D. Maystre, L. C. Botten, C. M. de Sterke, and R. C. McPhedran, “Multipole method for microstructured optical fibers. II. Implementation and results,” J. Opt. Soc. Am. B 19(10), 2331–2340 (2002). [CrossRef]  

28. T. Grujic, B. T. Kuhlmey, A. Argyros, S. Coen, and C. M. de Sterke, “Solid-core fiber with ultra-wide bandwidth transmission window due to inhibited coupling,” Opt. Express 18(25), 25556–25566 (2010). [CrossRef]   [PubMed]  

29. J. Jensen, P. Hoiby, G. Emiliyanov, O. Bang, L. Pedersen, and A. Bjarklev, “Selective detection of antibodies in microstructured polymer optical fibers,” Opt. Express 13(15), 5883–5889 (2005). [CrossRef]   [PubMed]  

30. R. Reiter, H. Motschmann, and W. Knoll, “Ellipsometric characterization of streptavidin binding to biotin-functionalized lipid monolayers at the water/air interface,” Langmuir 9(9), 2430–2435 (1993). [CrossRef]  

31. 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]  

32. H. Tazawa, T. Kanie, and M. Katayama, “Fiber-optic coupler based refractive index sensor and its application to biosensing,” Appl. Phys. Lett. 91(11), 113901 (2007). [CrossRef]  

33. Y. Zhang, H. Shibru, K. L. Cooper, and A. Wang, “Miniature fiber-optic multicavity Fabry-Perot interferometric biosensor,” Opt. Lett. 30(9), 1021–1023 (2005). [CrossRef]   [PubMed]  

34. M. Golic, K. Walsh, and P. Lawson, “Short-wavelength near-infrared spectra of sucrose, glucose, and fructose with respect to sugar concentration and temperature,” Appl. Spectrosc. 57(2), 139–145 (2003). [CrossRef]   [PubMed]  

References

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  1. S. Yin, P. B. Ruffin, and F. T. S. Yu, Fiber Optic Sensors, 2nd ed. (CRC Press, 2008).
  2. A. Hassani and M. Skorobogatiy, “Design criteria for microstructured-optical-fiber-based surface-plasmon-resonance sensors,” J. Opt. Soc. Am. B 24(6), 1423–1429 (2007).
    [Crossref]
  3. A. Hassani and M. Skorobogatiy, “Design of the microstructured optical fiber-based surface plasmon resonance sensors with enhanced microfluidics,” Opt. Express 14(24), 11616–11621 (2006).
    [Crossref] [PubMed]
  4. A. Wang, A. Docherty, B. T. Kuhlmey, F. M. Cox, and M. C. J. Large, “Side-hole fiber sensor based on surface plasmon resonance,” Opt. Lett. 34(24), 3890–3892 (2009).
    [Crossref] [PubMed]
  5. V. Bhatia and A. M. Vengsarkar, “Optical fiber long-period grating sensors,” Opt. Lett. 21(9), 692–694 (1996).
    [Crossref] [PubMed]
  6. L. Rindorf and O. Bang, “Highly sensitive refractometer with a photonic-crystal-fiber long-period grating,” Opt. Lett. 33(6), 563–565 (2008).
    [Crossref] [PubMed]
  7. D. K. C. Wu, B. T. Kuhlmey, and B. J. Eggleton, “Ultrasensitive photonic crystal fiber refractive index sensor,” Opt. Lett. 34(3), 322–324 (2009).
    [Crossref] [PubMed]
  8. G. E. Town, W. Yuan, R. McCosker, and O. Bang, “Microstructured optical fiber refractive index sensor,” Opt. Lett. 35(6), 856–858 (2010).
    [Crossref] [PubMed]
  9. W. Yuan, G. E. Town, and O. Bang, “Refractive index sensing in an all-solid twin-core photonic bandgap fiber,” IEEE Sens. J. 10(7), 1192–1199 (2010).
    [Crossref]
  10. Y. Zhang, H. Shibru, K. L. Cooper, and A. Wang, “Miniature fiber-optic multicavity Fabry-Perot interferometric biosensor,” Opt. Lett. 30(9), 1021–1023 (2005).
    [Crossref] [PubMed]
  11. Y. Gong, T. Zhao, Y.-J. Rao, Y. Wu, and Y. Guo, “A ray-transfer-matrix model for hybrid fiber Fabry-Perot sensor based on graded-index multimode fiber,” Opt. Express 18(15), 15844–15852 (2010).
    [Crossref] [PubMed]
  12. H. W. Lee, M. A. Schmidt, P. Uebel, H. Tyagi, N. Y. Joly, M. Scharrer, and P. St. J. Russell, “Optofluidic refractive-index sensor in step-index fiber with parallel hollow micro-channel,” Opt. Express 19(9), 8200–8207 (2011).
    [Crossref] [PubMed]
  13. D. K. C. Wu, K. J. Lee, V. Pureur, and B. T. Kuhlmey, “Performance of Refractive Index Sensors Based On Directional Couplers in Photonic Crystal Fibers,” J. Lightwave Technol. 31(22), 3500–3510 (2013).
    [Crossref]
  14. B. Sun, M.-Y. Chen, Y.-K. Zhang, J.-C. Yang, J.-Q. Yao, and H.-X. Cui, “Microstructured-core photonic-crystal fiber for ultra-sensitive refractive index sensing,” Opt. Express 19(5), 4091–4100 (2011).
    [Crossref] [PubMed]
  15. 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]
  16. L. Rindorf and O. Bang, “Highly sensitive refractometer with a photonic-crystal-fiber long-period grating,” Opt. Lett. 33(6), 563–565 (2008).
    [Crossref] [PubMed]
  17. M. Yang, D. N. Wang, Y. Wang, and C. Liao, “Long period fiber grating formed by periodically structured microholes in all-solid photonic bandgap fiber,” Opt. Express 18(3), 2183–2189 (2010).
    [Crossref] [PubMed]
  18. C. Wang and C. Herath, “High-sensitivity fiber-loop ringdown evanescent-field index sensors using single-mode fiber,” Opt. Lett. 35(10), 1629–1631 (2010).
    [Crossref] [PubMed]
  19. M. H. Frosz, A. Stefani, and O. Bang, “Highly sensitive and simple method for refractive index sensing of liquids in microstructured optical fibers using four-wave mixing,” Opt. Express 19(11), 10471–10484 (2011).
    [Crossref] [PubMed]
  20. C. Markos, W. Yuan, K. Vlachos, G. E. Town, and O. Bang, “Label-free biosensing with high sensitivity in dual-core microstructured polymer optical fibers,” Opt. Express 19(8), 7790–7798 (2011).
    [Crossref] [PubMed]
  21. B. T. Kuhlmey, S. Coen, and S. Mahmoodian, “Coated photonic bandgap fibres for low-index sensing applications: cutoff analysis,” Opt. Express 17(18), 16306–16321 (2009).
    [Crossref] [PubMed]
  22. S. G. Leon-Saval, R. Lwin, and A. Argyros, “Multicore composite single-mode polymer fiber,” Opt. Express 20(1), 141–148 (2012).
    [Crossref] [PubMed]
  23. A. Argyros, “Microstructured polymer optical fibers,” J. Lightwave Technol. 27(11), 1571–1579 (2009).
    [Crossref]
  24. http://www.zeonex.com/datasheets.asp
  25. B. T. Kuhlmey, B. J. Eggleton, and D. K. C. Wu, “Fluid-filled solid-core photonic bandgap fibers,” J. Lightwave Technol. 27(11), 1617–1630 (2009).
    [Crossref]
  26. I. M. White and X. D. Fan, “On the performance quantification of resonant refractive index sensors,” Opt. Express 16(2), 1020–1028 (2008).
    [Crossref] [PubMed]
  27. B. T. Kuhlmey, T. P. White, G. Renversez, D. Maystre, L. C. Botten, C. M. de Sterke, and R. C. McPhedran, “Multipole method for microstructured optical fibers. II. Implementation and results,” J. Opt. Soc. Am. B 19(10), 2331–2340 (2002).
    [Crossref]
  28. T. Grujic, B. T. Kuhlmey, A. Argyros, S. Coen, and C. M. de Sterke, “Solid-core fiber with ultra-wide bandwidth transmission window due to inhibited coupling,” Opt. Express 18(25), 25556–25566 (2010).
    [Crossref] [PubMed]
  29. J. Jensen, P. Hoiby, G. Emiliyanov, O. Bang, L. Pedersen, and A. Bjarklev, “Selective detection of antibodies in microstructured polymer optical fibers,” Opt. Express 13(15), 5883–5889 (2005).
    [Crossref] [PubMed]
  30. R. Reiter, H. Motschmann, and W. Knoll, “Ellipsometric characterization of streptavidin binding to biotin-functionalized lipid monolayers at the water/air interface,” Langmuir 9(9), 2430–2435 (1993).
    [Crossref]
  31. 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]
  32. H. Tazawa, T. Kanie, and M. Katayama, “Fiber-optic coupler based refractive index sensor and its application to biosensing,” Appl. Phys. Lett. 91(11), 113901 (2007).
    [Crossref]
  33. Y. Zhang, H. Shibru, K. L. Cooper, and A. Wang, “Miniature fiber-optic multicavity Fabry-Perot interferometric biosensor,” Opt. Lett. 30(9), 1021–1023 (2005).
    [Crossref] [PubMed]
  34. M. Golic, K. Walsh, and P. Lawson, “Short-wavelength near-infrared spectra of sucrose, glucose, and fructose with respect to sugar concentration and temperature,” Appl. Spectrosc. 57(2), 139–145 (2003).
    [Crossref] [PubMed]

2013 (1)

2012 (1)

2011 (4)

2010 (6)

2009 (5)

2008 (4)

2007 (2)

H. Tazawa, T. Kanie, and M. Katayama, “Fiber-optic coupler based refractive index sensor and its application to biosensing,” Appl. Phys. Lett. 91(11), 113901 (2007).
[Crossref]

A. Hassani and M. Skorobogatiy, “Design criteria for microstructured-optical-fiber-based surface-plasmon-resonance sensors,” J. Opt. Soc. Am. B 24(6), 1423–1429 (2007).
[Crossref]

2006 (1)

2005 (3)

2003 (1)

2002 (1)

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]

1996 (1)

1993 (1)

R. Reiter, H. Motschmann, and W. Knoll, “Ellipsometric characterization of streptavidin binding to biotin-functionalized lipid monolayers at the water/air interface,” Langmuir 9(9), 2430–2435 (1993).
[Crossref]

Argyros, A.

Bang, O.

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]

Bhatia, V.

Bjarklev, A.

Botten, L. C.

Chen, M.-Y.

Coen, S.

Cooper, K. L.

Cox, F. M.

Cui, H.-X.

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]

de Sterke, C. M.

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]

Docherty, A.

Eggleton, B. J.

Emiliyanov, G.

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]

Fan, X. D.

Frosz, M. H.

Golic, M.

Gong, Y.

Grujic, T.

Guo, Y.

Hassani, A.

Herath, C.

Hoiby, P.

Jensen, J.

Joly, N. Y.

Kanie, T.

H. Tazawa, T. Kanie, and M. Katayama, “Fiber-optic coupler based refractive index sensor and its application to biosensing,” Appl. Phys. Lett. 91(11), 113901 (2007).
[Crossref]

Katayama, M.

H. Tazawa, T. Kanie, and M. Katayama, “Fiber-optic coupler based refractive index sensor and its application to biosensing,” Appl. Phys. Lett. 91(11), 113901 (2007).
[Crossref]

Knoll, W.

R. Reiter, H. Motschmann, and W. Knoll, “Ellipsometric characterization of streptavidin binding to biotin-functionalized lipid monolayers at the water/air interface,” Langmuir 9(9), 2430–2435 (1993).
[Crossref]

Kuhlmey, B. T.

Large, M. C. J.

Lawson, P.

Lee, H. W.

Lee, K. J.

Leon-Saval, S. G.

Liao, C.

Lwin, R.

Mahmoodian, S.

Markos, C.

Maystre, D.

McCosker, R.

McPhedran, R. C.

Motschmann, H.

R. Reiter, H. Motschmann, and W. Knoll, “Ellipsometric characterization of streptavidin binding to biotin-functionalized lipid monolayers at the water/air interface,” Langmuir 9(9), 2430–2435 (1993).
[Crossref]

Pedersen, L.

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]

Pureur, V.

Rao, Y.-J.

Reiter, R.

R. Reiter, H. Motschmann, and W. Knoll, “Ellipsometric characterization of streptavidin binding to biotin-functionalized lipid monolayers at the water/air interface,” Langmuir 9(9), 2430–2435 (1993).
[Crossref]

Renversez, G.

Rindorf, L.

Russell, P. St. J.

Scharrer, M.

Schmidt, M. A.

Shibru, H.

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]

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]

Skorobogatiy, M.

Stefani, A.

Sun, B.

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]

Tazawa, H.

H. Tazawa, T. Kanie, and M. Katayama, “Fiber-optic coupler based refractive index sensor and its application to biosensing,” Appl. Phys. Lett. 91(11), 113901 (2007).
[Crossref]

Town, G. E.

Tyagi, H.

Uebel, P.

Vengsarkar, A. M.

Vlachos, K.

Walsh, K.

Wang, A.

Wang, C.

Wang, D. N.

Wang, Y.

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]

I. M. White and X. D. Fan, “On the performance quantification of resonant refractive index sensors,” Opt. Express 16(2), 1020–1028 (2008).
[Crossref] [PubMed]

White, T. P.

Wu, D. K. C.

Wu, Y.

Yang, J.-C.

Yang, M.

Yao, J.-Q.

Yuan, W.

Zhang, Y.

Zhang, Y.-K.

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]

Zhao, T.

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]

Anal. Chem. (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]

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. (1)

H. Tazawa, T. Kanie, and M. Katayama, “Fiber-optic coupler based refractive index sensor and its application to biosensing,” Appl. Phys. Lett. 91(11), 113901 (2007).
[Crossref]

Appl. Spectrosc. (1)

IEEE Sens. J. (1)

W. Yuan, G. E. Town, and O. Bang, “Refractive index sensing in an all-solid twin-core photonic bandgap fiber,” IEEE Sens. J. 10(7), 1192–1199 (2010).
[Crossref]

J. Lightwave Technol. (3)

J. Opt. Soc. Am. B (2)

Langmuir (1)

R. Reiter, H. Motschmann, and W. Knoll, “Ellipsometric characterization of streptavidin binding to biotin-functionalized lipid monolayers at the water/air interface,” Langmuir 9(9), 2430–2435 (1993).
[Crossref]

Opt. Express (12)

T. Grujic, B. T. Kuhlmey, A. Argyros, S. Coen, and C. M. de Sterke, “Solid-core fiber with ultra-wide bandwidth transmission window due to inhibited coupling,” Opt. Express 18(25), 25556–25566 (2010).
[Crossref] [PubMed]

J. Jensen, P. Hoiby, G. Emiliyanov, O. Bang, L. Pedersen, and A. Bjarklev, “Selective detection of antibodies in microstructured polymer optical fibers,” Opt. Express 13(15), 5883–5889 (2005).
[Crossref] [PubMed]

I. M. White and X. D. Fan, “On the performance quantification of resonant refractive index sensors,” Opt. Express 16(2), 1020–1028 (2008).
[Crossref] [PubMed]

M. Yang, D. N. Wang, Y. Wang, and C. Liao, “Long period fiber grating formed by periodically structured microholes in all-solid photonic bandgap fiber,” Opt. Express 18(3), 2183–2189 (2010).
[Crossref] [PubMed]

A. Hassani and M. Skorobogatiy, “Design of the microstructured optical fiber-based surface plasmon resonance sensors with enhanced microfluidics,” Opt. Express 14(24), 11616–11621 (2006).
[Crossref] [PubMed]

Y. Gong, T. Zhao, Y.-J. Rao, Y. Wu, and Y. Guo, “A ray-transfer-matrix model for hybrid fiber Fabry-Perot sensor based on graded-index multimode fiber,” Opt. Express 18(15), 15844–15852 (2010).
[Crossref] [PubMed]

H. W. Lee, M. A. Schmidt, P. Uebel, H. Tyagi, N. Y. Joly, M. Scharrer, and P. St. J. Russell, “Optofluidic refractive-index sensor in step-index fiber with parallel hollow micro-channel,” Opt. Express 19(9), 8200–8207 (2011).
[Crossref] [PubMed]

B. Sun, M.-Y. Chen, Y.-K. Zhang, J.-C. Yang, J.-Q. Yao, and H.-X. Cui, “Microstructured-core photonic-crystal fiber for ultra-sensitive refractive index sensing,” Opt. Express 19(5), 4091–4100 (2011).
[Crossref] [PubMed]

M. H. Frosz, A. Stefani, and O. Bang, “Highly sensitive and simple method for refractive index sensing of liquids in microstructured optical fibers using four-wave mixing,” Opt. Express 19(11), 10471–10484 (2011).
[Crossref] [PubMed]

C. Markos, W. Yuan, K. Vlachos, G. E. Town, and O. Bang, “Label-free biosensing with high sensitivity in dual-core microstructured polymer optical fibers,” Opt. Express 19(8), 7790–7798 (2011).
[Crossref] [PubMed]

B. T. Kuhlmey, S. Coen, and S. Mahmoodian, “Coated photonic bandgap fibres for low-index sensing applications: cutoff analysis,” Opt. Express 17(18), 16306–16321 (2009).
[Crossref] [PubMed]

S. G. Leon-Saval, R. Lwin, and A. Argyros, “Multicore composite single-mode polymer fiber,” Opt. Express 20(1), 141–148 (2012).
[Crossref] [PubMed]

Opt. Lett. (9)

L. Rindorf and O. Bang, “Highly sensitive refractometer with a photonic-crystal-fiber long-period grating,” Opt. Lett. 33(6), 563–565 (2008).
[Crossref] [PubMed]

Y. Zhang, H. Shibru, K. L. Cooper, and A. Wang, “Miniature fiber-optic multicavity Fabry-Perot interferometric biosensor,” Opt. Lett. 30(9), 1021–1023 (2005).
[Crossref] [PubMed]

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Other (2)

S. Yin, P. B. Ruffin, and F. T. S. Yu, Fiber Optic Sensors, 2nd ed. (CRC Press, 2008).

http://www.zeonex.com/datasheets.asp

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Figures (6)

Fig. 1
Fig. 1 (a) Schematic diagram of the directional coupler geometry within a polymer fiber. (b) Details of the cross section; (c) Schematic of effective index curves of core and satellite modes and resonance wavelength.
Fig. 2
Fig. 2 (a) Schematic design of the composite core. The intensity profiles of the lowest order modes in (b) composite and circular single-mode cores using (c) Zeonex and (d) PC. All cases shown support two degenerate modes of orthogonal polarization: LP01x and LP01y (or HE11x and HE11y).
Fig. 3
Fig. 3 (a) Schematic design of the satellite waveguide (rin = 2 μm, t = 300 nm). The intensity profiles of four guided modes in the satellite waveguide when the hole coated with a PC ring is filled with low-index analyte (na = 1.333): (b) TM, (c) TE, (d) HE-1 (e) HE-2 modes.
Fig. 4
Fig. 4 (a) Calculated effective index for the core and four satellite modes showing the resonant wavelength, and (b) the coupling length for HE11x-TM mode combination plotted as a function of the center-to-center distance (h) between the core and satellite waveguides. The distance, h, is defined in the inset of Fig. 4(b).
Fig. 5
Fig. 5 (a) Calculated modal intensity overlap in the analyte hole (fsat) for TM and TE modes, and (b) fsat curves for TM mode plotted for different sizes of the analyte hole.
Fig. 6
Fig. 6 (a) Microscope image of the core and analyte channel of the fiber; (b) transmission spectra through the empty fiber (blue), and after infiltration for various concentration of sucrose, as indicated – spectra are offset for readability; (c) Measured wavelength of the dips indicated by red arrows in (b) as a function of refractive index. The error bars reflect the accuracy of the refractometer ( ± 0.0003) and of the OSA ( ± 0.1 nm).

Tables (3)

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Table 1 Coupling lengths for each core-satellite mode combination at the corresponding resonant wavelengths, for h = 10 μm. Only one polarization for each degenerate pair shown.

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Table 2 Detection limits for fiber-based low-index sensing reported in recent publications.

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Table 3 Detection limits for fiber-based protein measurement reported in recent publications.

Equations (2)

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δnα λ r L c f sat T min (SNR) 0.25 ,
f sat = analyte ε sat (x,y) | E sat (x,y) | 2 dxdy cross secsion ε sat (x,y) | E sat (x,y) | 2 dxdy ,

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