Fiber light-coupled optofluidic waveguide (FLOW) immunosensor for

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Fiber light-coupled optofluidic waveguide (FLOW) immunosensor for highly sensitive detection of p53 protein Lili Liang, Long Jin, Yang Ran, Li-Peng Sun, and Bai-Ou Guan Anal. Chem., Just Accepted Manuscript • DOI: 10.1021/acs.analchem.8b02123 • Publication Date (Web): 24 Aug 2018 Downloaded from http://pubs.acs.org on August 27, 2018

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Analytical Chemistry

Fiber light-coupled optofluidic waveguide (FLOW) immunosensor for highly sensitive detection of p53 protein †

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Lili Liang , Long Jin*, , Yang Ran , , Li-Peng Sun , Bai-Ou Guan

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Guangdong Provincial Key Laboratory of Optical Fiber Sensing and Communication, Institute of Photonics Technology, Jinan University, Guangzhou 510632, China ‡

Department of Biomedical Engineering, Duke University, Durham, 27708, United States

ABSTRACT: Highly sensitive detection of molecular tumor markers is essential for biomarker-based cancer diagnostics. In this work, we showcase the implementation of Fiber Light-coupled Optofluidic Waveguide (FLOW) immunosensor for the detection of p53 protein, a typical tumor marker. The FLOW consists of a liquid-core capillary and an accompanying optical fiber, which allows evanescent interaction between light and microfluidic sample. Molecular binding at internal surface of the capillary induces a response in wavelength shift of the transmission spectrum in the optical fiber. In order to enable highly sensitive molecular detection, the evanescent-wave interaction has been strengthened by enlarging shape factor R via fine geometry control. The proposed FLOW immunosensor works with flowing microfluid, which increases the surface molecular coverage and improves the detection limit. As a result, the FLOW immunosensor presents a log-linear response to the tumor protein at concentrations ranging from 10 fg/mL up to 10 ng/mL. In addition, the non-specifically adsorbed molecules can be effectively removed by the fluid at an optimal flow rate, which benefits the accuracy of the measurement. Tested in serum samples, the FLOW successfully maintains its sensitivity and specificity on p53 protein, making it suitable for diagnostics applications.

Biomarker-based diagnostics demands on fast, sensitive and preferably label-free detection of biomolecules which are indicative of the presence and development of cancer in human body. For instance, p53 protein (molecular weight: 53 kDa) is a typical tumor suppressor and transcription factor in cancer biology.1-3 It plays a crucial role in the regulation of the cell cycle, DNA repair, and programmed cell death, inhibiting the growth of tumor cells through eliciting either cell-cycle arrest or apoptosis.4 This protein is related to more than 50% human cancers 5 and its concentration in human blood can be reasonably linked with tissue alterations.6 Therefore, quantitatively detection of protein p53 can help early-stage cancer diagnostics and related risk assessment. 4,7,8 Detection of protein p53 has been achieved taking advantage of quartz crystal microbalance,9 electrochemical 4,10-13 immunosensors, chemiluminescence,5,14,15 as well as surface enhanced Raman scattering (SERS).16,17 Most methods involve additive signal amplification to attain sufficient sensitivity by adsorbing more target biomolecules at the binding sites, i. e., the functionalized surface or nanoparticle probes, which complex the detection procedures. Point-of-care diagnostics demands the development of biomarker assays with high sensitivity, specificity, low sample consumption, limited detection duration as well as field portability. Optofluidic immunosensors, which opti-

cally detect the biomarkers in microfluidics, have been developed towards biological analysis for medical applications. For example, surface plasmon resonance (SPR)18 sensors can detect the bound biomolecules at the metalcoated optical surface based on the strong evanescent wave interaction.19 Current portable SPR-based bioanalysis instrumentations measure the binding-induced spectral change.20 However, maximum sensitivity can be attained by prism-coupling the probe light exactly at the angle of total reflection and measuring the polarization-dependant intensity change of the reflected-off light. The performance of the SPR sensor has been limited by the trade-off between its output stability and sensitivity.21 Liquid-core waveguide sensors22-24 have been fabricated by composing high/low-index alternating dielectric layers (namely ARROW waveguide) to confine light in the liquid core, which allows a strong overlap between light and the liquid core and enables highly sensitive fluorescent detection.23,25,26 Alternatively, optical micro-resonators27-29 have been exploited as optofluidic sensors, utilizing the detection of whispering gallery modes (WGMs).30-32 The WGM sensors which drive interaction between light and target molecules by multiple times,33,34 can detect biomolecules at extremely low concentration,35 even to singlemolecule level.36-38 For example, the liquid-core opticalring-resonator (LCORR) sensor integrated an optical resonator and microfluidics, which offers a convenient plat-

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form for microfluidic bioanalysis. 39,40 By integrating multiple micro-resonators on a silicon photonic chip, a multiplexed sensor array can be implemented for simultaneous measurement of different target molecules. 41-43 Design of an optofluidic sensor44 for point-of-care applications mainly involves the following two aspects. First, the sensor should be sufficiently sensitive to detect the biomarkers at extremely low concentrations. Taking p53 quantification for example, cancer diagnostics requires a detection limit at as low as pg/mL level.6,45 The detection capability relies on how many molecules are bound to the functionalized surface with a given bulk concentration, as well as how a binding event is translated into optical signals. Second, the light path and the microfluidics should be incorporated in a highly integrated manner.46 With inherent connection to microfluidics, additive introduction of fluid inlet/outlet is not needed and the sensor fabrication is less sophisticated. Our group is particularly interested in how to develop such an optofluidic sensor based on low-cost fiber optics and related processing techniques such as thermal tapering. Based on our proof-of-concept demonstration in Ref 8, fiber lightcoupled optofluidic waveguide (FLOW) immunosensor is presented here for highly sensitive detection of p53 protein. The FLOW immunosensor presents a response from 10 fg/mL to 10 ng/mL and the serum test shows well maintained detection ability and specificity. That means the present FLOW sensor can respond to slight change in biomarker concentration in human serum, towards diagnostics applications. The immunosensor also has improved portability and high resistance to environmental perturbation. These advantages make it a suitable candidate for point-of-care application.

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ing processes). Figure 1b shows the transverse geometry at the taper waist, i. e., the optofluidic waveguide, which allows light/microfluidics overlapping, as schematically exhibited in Figure 1c. Its transverse geometry is characterized by the following parameters: capillary outer diameter D= 26.6 µm, wall thickness t=2 µm, fiber diameter d=5.6 µm. The optofluidic waveguide has a length of 2 cm. Broadband light from a source (Golight, OS-EB-S-D-1450-400-30-0-FA) is coupled to the optofluidic waveguide via the lead-in fiber and propagates in terms of two discrete optical modes, individually marked as “fiber mode” and “wall mode”. These two modes interfere with each other and form an interference spectrum, measured at the distal end of the optical fiber by using an optical spectrum analyzer (OSA, Yokogawa, AQ6370C). The tapered structure is embedded in the UV curable polymer for stability improvement.

SENSOR FABRICATION AND CHARACTERIZATION

Figure 1. (a) Schematic of the FLOW immunosensor. (b) SEM image of the transverse geometry of the optofluidic waveguide. (c) Schematic of the interaction between light and bound molecules at the capillary interior.

Figure 1a shows the schematic of the FLOW immunosensor. The capillary acts as the microfluidic channel. The optical fiber is used to illuminate the sensor and detect the optical response. They are parallel aligned and simultaneously tapered to form a dumbbell shape (See Figure S1 for details of taper-

Figure 2. Calculated intensity profiles of the two optical modes, named “fiber mode” and “wall mode”, of (a) present optofluidic waveguide with t/d=0.36 and (b) contrast waveguide with t/d=0.32. (c) Comparison on optical intensities at the internal surface of capillary corresponding to the blue dashed curves in (a) and (b) at the azimuthal direction, which indicates a significant enhancement in evanescent-field using the Present waveguide.

The periodic transmission spectrum is a result of linear variation of phase difference between the two modes ∆φ with optical wavelength. Biomolecular binding can change the local refractive index at the capillary interior, induce phase changes of both the modes, and shift the transmission spectrum. The sensitivity is proportion to the biomolecular surface density, as well as the difference in evanescent-field strength between the two modes. Based on the Langmuir adsorption


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Analytical Chemistry isotherm, 47,48 the adsorption rate of biomolecules at the functional surface reduces with increasing bulk concentration, especially at close to saturation. That means a given noise level can cause a higher measurement inaccuracy at high concentration. For optofluidic biosensors, an effective solution is to strengthen the evanescent-wave interaction. We found that the evanescent-field profiles are mainly determined by the transverse geometry of the optofluidic waveguide, characterized by a shape factor R = t/d, where t denotes capillary wall thickness and d fiber diameter. Figure 2a exhibits the intensity profiles of the two optical modes for the present optofluidic waveguide with R = 0.36, numerically calculated by solving the eigenmodes of the given waveguide. Figure 2b shows the calculation result of the proof-of-concept demonstration with R = 0.32 (d=6.9 µm, t=2.2 µm, and D= 32.5 µm) in Ref 8. With higher R, the optical modes tend to shift their mode energies towards the liquid core, yielding an intensity enhancement in the fluid sample. Figure 2c shows that the evanescentfiled strength at the capillary interior has been enhanced by a factor 1.98. Therefore, a sensitivity enhancement for the detection of bound molecules has been achieved using the Present waveguide. In order to achieve higher R, we attempt to increase t by slowly elongating the whole taper region, which allows the molten capillary to shrink further as a result of surface tension. Nevertheless, the sensitivity can’t be unlimitedly improved. Further enhancement of this shape factor R would cause reduced contrast in the interference spectrum, as a result of the unbalanced intensity redistribution between the two modes.

Protein P53 Detection in PBS Buffer. Figure 3a and Figure 3b illustrate the procedure for p53 protein detection in PBS buffer, which contains the following treatments to the interior of the capillary: (1) cleaning with piranha solution for 30 minutes; (2) modification with 5% APTES for 1 h to form amidogen (-NH2) on the surface; (3) immobilization of activated p53 antibody for 1h; (4) incubation with p53 protein solution with a certain bulk concentration for 1h after BSA blocking. The capillary is rinsed with PBS for 10 min to remove unbound or loosely bound molecules after each treatment. The optical repeating the above procedures. The experimental setup for p53 protein detection is shown in Figure S2. Specificity Test. For specificity test, we selected HAS,49 BSA,50 AFP,51 IgG52 and PSA5 as control proteins, which are serum proteins or tumor biomarkers in blood but having different molecular structures and functions. The optical responses to the individual control proteins are measured by the FLOW sensor functionalized by anti-p53 antibody. A blank control trial was also performed for contrast.

P53 PROTEIN DETECTION Materials. 98% sulfuric acid, hydrogen peroxide, phosphate buffered saline (PBS), absolute ethanol, antibody diluent (10×), tween20, bull serum albumin (BSA), Blocking Buffer BSA, anti-p53 antibody, p53 protein, human serum albumin(HSA) antigen, akpha-fetoprotein (AFP) antigen and immunoglobulin G (IgG) antigen were obtained from Sangon Biotech (Shanghai, China). 99% 3aminopropyltrimethoxysilane (APTES) was purchased from Sigma-Aldrich (Darmstadt, Germany). Prostate specific antigen (PSA) was obtained from Cholun Medical (Shenzhen, China). 1-(3-Dimethylaminopropyl) -3-ethylcarbodiimide hydrochloride (EDC) and N-Hydroxysuccinimide (NHS) were purchased from Aladdin (Shanghai, China). Human serum was obtained from Biopanda (Huzhou, China). DNA probe, modified with Cy3 fluorescent, and its complementary sequences were obtained from Sangon Biotech (Shanghai, China). The DNA probe had the sequences: 5’-Cy3-TTT TTA GAC ATG CCC AGA CAT GCC-3’ and its complementary DNA sequences are: 5’-GGG CAT GTC TGG GCA TGT CT-3’. Piranha solution was prepared by mixing the 98% sulfuric acid and hydrogen peroxide with a volume ratio 3:1. 5% APTES was prepared by dissolving 99% APTES in absolute ethanol. EDC/NHS solution was prepared by dissolving 0.23 g NHS in 500 mL distilled water and then adding 0.16 g EDC and 500 mL deionized water. Anti-p53 antibody of 10 ng/mL was prepared using antibody diluent and activated by EDC/NHS with ratio 100:3 in swing bed for two hours before using. P53 antigen was diluted using PBS to 10 ng/mL, 1 ng/mL, 100 pg/mL, 10 pg/mL, 1 pg/mL, 100 fg/mL, 10 fg/mL, 1 fg/mL. P53 antigen was also diluted using human serum to 1 ng/mL and 100 pg/mL. HSA, IgG, BSA, AFP and PSA were diluted using PBS to 10 ng/mL respectively.

Figure 3. Schematic diagram(a), procedure chart (b) and AFM images (c), which depict and characterize the functionalization and p53 protein detection procedures, including (1) cleaning with piranha solution; (2) APTES modification; (3) p53 antibody immobilization; (4) p53 protein incubation. AFM images are taken after each step is finished. (d) Fluorescent image of microcapillary before (left) and after (right) p53 protein incubation after infiltration of fluorescent-labeled DNA probe.

Protein P53 Detection in Serum. For most cancer patients, the concentration of p53 protein increases significantly in human blood,5 typically ranging from 100 pg/mL to 1 ng/mL. Therefore, we have prepared p53 protein serum samples with 100 pg/mL and 1 ng/mL to test the validity of the FLOW immunosensor. The samples were prepared by adding p53 antigens to the serum without any pre-treatment for demonstrating the detection under a more complex and realistic environment53 Based on the same functionalization procedures


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as shown in Figure 3a and Figure 3b, the optical responses to serum samples with different concentrations were measured. On the other hand, the serum samples are also measured by use of a contrast sensor, fabricated by straightforward tapering a single mode fiber. The tapered sensor is also a modal interferometer (See Figure S3 in Supporting materials for more details),54,55 in analogy to the present optofluidic waveguide sensor. The evanescent-field strength is close to the FLOW sensor, for better comparison. The tapered sensor was consecutively immersed into the solutions for functionalization and incubation of target molecules.53,56

RESULTS AND DISCUSSION

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shows measured responses to p53 proteins at different concentrations. Each error bar represents the standard deviation in three consecutive independent measurements (including the functionalization and binding procedures) with the same sensor. The spectrum red-shifted almost log-linearly with the concentration ranging from 10 fg/mL to 10 ng/mL and began to level off with concentrations below 10 fg/mL. The linear regression equation is expressed as y (nm) =0.32+0.24 × log (C (fg/mL)) (R2 = 0.99). When the bulk concentration approaches zero, the fitted sensitivity is about S=22.2 pm/(fg/mL). We have measured the wavelength drift of the FLOW sensor, by recording the spectrum for 150 min after p53 protein incubation. The wavelength fluctuation is shown in the inset of Figure 4b. The standard deviation σ =30 pm, which is mainly induced by thermal noise. The theoretical limit of detection can be estimated as LOD=3σ/S= 4 fg/mL.57,58 In practical measurement, however, the fluid flow and internal pressure induced by the mechanical pump causes additional noise, which limits the measurement capability approaching to the theoretical limit. Nevertheless, the experimentally confirmed LOD at 10 fg/mL to p53 proteins is sufficient for diagnostic applications.

Figure 4. (a) Sensorgram for surface activation and p53 detection at a bulk concentration of 10 pg/mL. Inset: Recorded spectral shift induced by p53 bimolecular binding. (b) Measured response as a function of p53 concentration. Inset: Measured dip wavelength drift over 150 min. Standard wavelength deviation σ=30 pm.

P53 Protein Detection in PBS Buffer. Prior to biomolecular detection, the validity of each procedure has been verified from the roughness of the fiber surface, as shown by the AFM images in Figure 3c. To verify the p53 protein adsorption, immuno-fluorescence assays were carried out after p53 protein incubation in a micro-capillary by reacting with DNA probe modified with fluorescent group Cy3, which specifically adsorb p53 protein. The fluorescence images in Figure 3d verifies that p53 protein has specifically bound to the functionalized surface of the capillary. Figure 4a shows the sensorgram for procedures 2 to 4, to measure the response to p53 protein at a bulk concentration 10 pg/mL. The transmission spectrum was measured every 30 seconds with a fluid velocity 3.7 × 1011 m3/s. The binding induced spectral shift is 1.2 nm. Figure 4b

Figure 5. (a) Measured wavelength responses to different tumor biomarkers and PBS, comparing the target analyte p53 protein. (b) Wavelength responses of optofluidic sensor to 100 pg/mL and 1 ng/mL of target analyte p53 protein in PBS and serum conditions.

Specificity. Figure 5a shows the measured responses to control proteins and PBS, in comparison to p53 protein result. The test samples have identical concentrations at 10 ng/mL in PBS. The result shows that the target biomolecule can induce much more wavelength shift than others. The stronger re-


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Analytical Chemistry sponse in wavelength shift suggest good specificity of the sensor for p53 protein recognition in PBS condition. P53 Protein Detection in Serum. Figure 5b shows the result of serum test, compared to PBS test. The optical responses are 1.44 and 1.81 nm, according to serum samples at concentrations 100 pg/mL and 1 ng/mL respectively. These responses are almost identical to that for PBS samples with 1.48 and 1.79 nm. Different from the PBS solution, the serum samples contain sera proteins, other proteins and biomolecules, which can induce nonspecific adsorption and possibly reduce the p53-protein binding probability (See Figure S4 for the comparative results of the tapered fiber sensor).51 In the FLOW immunosensor, optimal fluidic flow velocity can effectively minimize nonspecific serum proteins adsorption and maintain the specific binding (to p53 protein) probability, taking advantage of the difference in molecular binding strength. This will be discussed in the next section. In contrast, Serum without p53 proteins can enable only 0.2 nm wavelength shift as a result of non-specific adsorption. Discussion on the Effect of Fluid Flow. The present FLOW immunosensor works with flowing fluid, different to other fiber-coupled biosensors immersed into static solutions. This section discusses the effect of the fluid flow on the detection capability. For the FLOW sensor in static mode, the reaction kinetics between the biological probe anti-p53 antibody and p53 protein analyte can be regarded as an equilibrium process of adsorption and desorption.59 The reaction kinetics can be described by the isothermal Langmuir adsorption model ⁄ (1) where is the total concentration of antibody sites on the surface, the surface concentration of antibodies that are bound by antigens at any time, denotes antigen concentration near the surface, and are the forward and reverse reaction-rate constants. Figure 6 shows the schematic of the binding kinetics. At static mode, the surface adsorption rate of the targets gradually slows down and eventually reaches equilibrium.

where DL denotes the diffusion coefficient and represents the velocity vector. According to Eq 1, bound surface density of antibody after reaction equilibrium is increased with enhanced . At static mode, only a very limited number of molecules can be adsorbed, as a result of the slow process of free diffusion. In contrast, at the flow mode, the flow velocity accelerates the mass transfer and approximates the concentration near the surface to the bulk concentration . Based on Eq 1, the raised can effectively induce more molecules adsorbed on the functionalization surface. As a result, the flow-mode sensor can shift the static-mode response curve (Fig. 4b) horizontally towards lower bulk concentration, which would significantly lower the detection limit. The fluidic flow also generates a shear stress at the sensing surface, which can cause effective desorption of bound molecules. We experimentally examined the effect of the shear stress at different flow velocities. The capillary was functionalized and incubated with p53 protein in static mode, before PBS rinsing step with increased velocities. Figure 7 shows that the unbound and non-specifically bound molecules are removed under velocities at 1.8×10-11, 3.7×10-11 and 5.2×10-11 m3/s.61 However, further increasing velocity to 6.7×10-11 m3/s, induces an additional optical wavelength change, as a result of partial removal of the specifically bound molecules,62 which is undesired for detection. Therefore, a moderate fluidic flow velocity helps to reduce non-specific absorption, and the flow velocity was kept below 5.2×10-11 m3/s for the FLOW immunosensor used in this experiment with geometries shown in Figure 1a.

Figure 7. Wavelength shift during PBS rinsing after p53 protein incubation with increasing velocities.

CONCLUSION

Figure 6. Schematic of binding kinetics in the static / flow mode

Notably, the amplitude of is also related to the bulk concentration through mass transfer process, which can be described by 60 ⁄ ∙ ∙ 0 (2)

In conclusion, we have demonstrated a fiber light-coupled optofluidic waveguide immunosensor for the detection of a typical tumor biomarker, p53 protein. The FLOW sensor has a log-linear response to the biomarkers at concentrations ranging from 10 fg/mL to 10 ng/mL. Different from the previous fiber optic biosensors, the FLOW immunosensor operates with flowing microfluidics, which makes a positive contribution to the sensitivity by changing the concentration contrast at the sensing surface, as well as specificity by desorbing the nonspecific molecules. As a result, serum test shows the sensitivity and specificity can be well maintained, which indicates the validity for further diagnostic applications. In addition, the FLOW sensor presents inherent resistance to environmental perturbation as a result of the highly integrated manner be-


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tween light and microfluidics, yielding improved field portability. We attempt to exploit the FLOW immunosensor as point-of-care testing instrumentations towards medical diagnostics.

ASSOCIATED CONTENT Supporting Information Fabrication of the FLOW immunosensor, experimental set up for protein p53 detection, characterization of the contrast immunosensor, and comparison with LCORR sensors.

AUTHOR INFORMATION Corresponding Author * E-mail: [email protected]. Notes The authors declare no competing financial interest.

ACKNOWLEDGMENT This work was supported by National Natural Science Foundation of China under Grants U1701268 and 61775082. Y. R. was supported by a grant from the China Scholarship Council (CSC No. 201706785006).

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Fiber light-coupled optofluidic waveguide (FLOW) immunosensor for

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