Letter Vol. 45, No. 24 / 15 December 2020 / Optics Letters 6595 Coherent silicon photonic interferometric biosensor with an inexpensive laser source for sensitive label-free immunoassays Jonas Leuermann,1,2, * Vladimir Stamenkovic,1,3 Patricia Ramirez-Priego,4 Alejandro Sánchez-Postigo,2 Adrián Fernández-Gavela,5 Cole A. Chapman,6 Ryan C. Bailey,6 Laura M. Lechuga,4 Ezequiel Perez-Inestrosa,1,3 Daniel Collado,1,3 Robert Halir,1,2 AND Íñigo Molina-Fernández1,2 1 Bionand Center for Nanomedicine and Biotechnology, Parque Tecnológico de Andalucía, 29590 Málaga, Spain 2 Departamento de Ingeniería de Comunicaciones, University of Málaga, ETSI Telecomunicación, Campus de Teatinos, 29071 Málaga, Spain 3 Departmento de Química Orgánica, Facultad de Ciencias, University of Málaga, 29071 Málaga, Spain 4 Nanobiosensors and Bioanalytical Applications Group, Catalan Institute of Nanoscience and Nanotechnology (ICN2), CSIC, BIST and CIBER-BBN Campus UAB, 08193 Barcelona, Spain 5 Departamento de Física, Universidad de Oviedo, C/Federico García Lorca, 33007 Oviedo, Spain 6 Department of Chemistry, University of Michigan, 930 North University Avenue, Ann Arbor, Michigan 48109, USA *Corresponding author: jonas.leuermann@uma.es Received 2 October 2020; revised 7 November 2020; accepted 9 November 2020; posted 10 November 2020 (Doc. ID 411635); published 8 December 2020 Over the past two decades, integrated photonic sensors personnel [5]. Most photonic integrated sensors rely on the have been of major interest to the optical biosensor com- principle of evanescent field sensing: the analyte attaches to a munity due to their capability to detect low concentrations biorecognition layer on the waveguide surface, where it interacts of molecules with label-free operation. Among these, with the evanescent field of the guided wave, changing its phase interferometric sensors can be read-out with simple, [6–8]. Since changes in the optical phase cannot be measured fixed-wavelength laser sources and offer excellent detec- directly, they are converted into changes in optical power, using tion limits but can suffer from sensitivity fading when on-chip resonant or interferometric structures. not tuned to their quadrature point. Recently, coherently Resonant structures convert the phase change into a res- detected sensors were demonstrated as an attractive alter- onance wavelength shift, thus often requiring a read-out native to overcome this limitation. Here we show, for the system with a tunable external cavity laser (ECL) [9–14], or a first time, to the best of our knowledge, that this coher- broadband light source combined with additional wavelength ent scheme provides sub-nanogram per milliliter limits filtering [15]. For homogeneous refractive index sensing, these of detection in C-reactive protein immunoassays and that systems offer moderate limits of detection (LODs), on the quasi-balanced optical arm lengths enable operation with order of 10−6 refractive index units (RIUs), but can achieve inexpensive Fabry–Perot-type lasers sources at telecom very compact footprints [9,16,17]. Regarding the detection wavelengths. © 2020 Optical Society of America under the terms of biomolecules, in benchmark primary C-reactive protein of the OSA Open Access Publishing Agreement (CRP) binding assays, resonant sensors achieve LODs on the https://doi.org/10.1364/OL.411635 order of 10 ng/mL (0.4 nM) [18]. Interferometric sensors compare the optical wave in the sensing waveguide to a wave Integrated photonic biosensors have been the subject of intense in a reference waveguide to reconstruct the phase shift. These research in the last two decades, due not only to their minituriz- sensors generally operate with a fixed-spectrum laser source, which can be a superluminescent diode, with asymmetric sens- ability but also because they can quantitatively detect extremely ing and reference arm and optical wavelength filtering in the low concentrations of analytes in real time and without the need read-out [2,19,20], or a narrow-linewidth laser source, e.g., a of labeling tags [1–3]. Silicon-based sensors provide particularly He–Ne laser at visible wavelengths [21–23] or ECLs at telecom compact footprints, enabling multiplexed detection of different wavelengths [10,24,25]. Detection limits down to 2ng/mL biochemical substances, and CMOS compatible fabrication, (84 pM) with a label-free primary binding immunoassay for which paves the way towards low-cost mass production [4]. detecting CRP have been reported for Mach–Zehnder inter- Therefore, they are candidates for point-of-care (POC) solu- ferometers (MZIs) [16,26]. It is noteworthy that shortcomings tions in which pre-functionalized, disposable chips are used in commonly associated with these interferometric sensors, such conjunction with a read-out system to provide diagnostics that as sensitivity fading when not operated at the quadrature point would otherwise require specialized laboratories and trained and phase ambiguity, have been overcome, either by adding a 0146-9592/20/246595-04 Journal © 2020 Optical Society of America 6596 Vol. 45, No. 24 / 15 December 2020 / Optics Letters Letter In our experiments, the reference and sensing arms have the same length L = 8.8 mm and a group index difference of 1n g,rs = |n g,s − n g,r | ≈0.01 at λ0 = 1.55 µm. In our read-out system, the light source is connected to an optional erbium-doped fiber amplifier and a variable optical attenuator to obtain the desired input power. A polarization rotator is used to achieve TE polarization that is coupled into the silicon nitride sensor chip via grating couplers, with coupling losses of ∼10 dB [Fig. 1(a)]. The sensor chip is fabricated with a multi-project wafer approach in a 300 nm thick silicon nitride film [36]. The sensing arm is covered with a polydimethyl- siloxane (PDMS) microfluidic channel, with a width of 1 mm, length of 9.8 mm, and height of 0.5 mm, which is connected Fig. 1. (a) Schematic of a coherently read balanced Mach–Zehnder to a syringe pump. In between, a six-port valve system enables sensor, where the phase difference between the reference and sensing the injection of an analyte volume of Vinj ≈120 µL into the arms is recovered from three output signals. The wavelength response channel. The three optical output signals are coupled to a of such a sensor is completely flat when the optical arm lengths are fiber array using on-chip identical grating couplers and are balanced. (b) Approximated laser spectra of a narrow-linewidth exter- detected by separate photodiodes (PDs) with integrated tran- nal cavity laser and a Fabry–Perot laser, with center wavelength λ0 , simpedance amplifiers. The resulting voltages are quantized linewidth δλ, and spectral spacing 1λ. by a data acquisition board, and signal processing is used to estimate the phase change 1ϕ [31]. A detailed description modulation signal [27–29] or by using a coherent detection and bulk LOD optimization of the experimental setup can be approach, illustrated in Fig. 1(a) [24,30–34]. With coherent found in [24]. To assess the impact of the laser linewidth on detection, no tuning to a quadrature point is required, so that sensor performance, two different lights sources were employed the optical lengths of the interferometer arms can be perfectly in the experiments. As a high-quality reference, we used the matched. So far, this approach has been demonstrated only with WSL-100 from SANTEC (Japan), a single-mode ECL bought bulk sensing experiments [10,24], and using high-quality ECLs for approximately 10,000€, with a linewidth of δλ < 80 fm with narrow linewidths [see Fig. 1(b)]. Here we show, for the (100 kHz) [Fig. 1(b)]. As a low-cost alternative, we used a first time, that this coherent approach is well suited for protein hand-held Fabry–Perot laser (FPL) with a center wavelength detection, reporting a limit of detection better than 300 pg/mL of λ0 = 1.543 µm, 1λ = 1.36 nm wavelength spacing, and a for primary binding CRP detection. We furthermore show bandwidth of δλ ∼ 3.6 nm (450 GHz) [Fig. 1(b)], purchased that due to quasi-identical optical arm lengths, the sensor can on Amazon for 34€. operate with a low-cost Fabry–Perot telecom laser [Fig. 1(b)] Two types of experiments were chosen to determine the with an acceptable performance penalty. quality of the sensor, bulk refractive index sensing, and pro- Mach–Zehnder-based sensors split the incoming light into tein CRP immunodetection. We used two different power two single-mode waveguides, one acting as the sensing arm and levels: low, with a total optical power of approximately of the other acting as the reference arm [Fig. 1(a)]. The reference −30 dBm received by the PDs, and high, with a power of waveguide is buried in silicon dioxide, so that the propagating approximately −10 dBm. Homogeneous refractive index sens- mode always experiences a constant phase shift. In sensing ing experiments were performed by injecting four different waveguides, the propagating mode interacts with the analyte, mass percentage solutions of sodium chloride (3%, 6%, 9%, which causes an additional phase shift. Conventional MZIs 12%) resulting in changes of (0.5, 1.0, 1.5, 2.0) × 10−2 RIU recombine these two arms into one output waveguide (e.g., using a Y-Junction or a 1 × 2 multi-mode interference cou- pler). This results in a signal proportional to sin(1ϕ), where 1ϕ is the phase difference between the waves in the two arms. This signal is affected by relative intensity noise, and causes sensitivity fading and phase ambiguity [16]. In contrast, a coher- ently read MZI recombines the sensing and reference arms via a 2 × 3 coupler [Fig. 1(a)]. From the three output powers, a complex signal ∝ exp( j 1ϕ) can be calculated [31,35], with the extracted phase being immune to relative intensity noise, sensitivity fading, and phase ambiguity [34]. The noise in the extracted√phase caused by a finite laser linewidth δ f is given by 21τ π δ f B, where B is the measurement bandwidth, and 1τ is the time delay between the reference and sensing arms, 1τ = |L s n g,s − L r n g,r |/c 0 , with L s , L r the physical lengths of the sensing and reference arms, n g,s , n g,r the group indices, and c 0 the speed of light [34]. Thus, by choosing L s Fig. 2. Bulk sensing using the reference ECL source (continuous and L r such that 1τ = 0, the phase noise can be minimized. blue curve) and the low-cost FPL (dashed orange curve) for 3%, 6%, Intuitively, this can be understood because for a perfectly bal- 9%, and 12% mass percentage sodium chloride injections and low anced interferometer, the wavelength response is flat [Fig. 1(a)]. input power. Letter Vol. 45, No. 24 / 15 December 2020 / Optics Letters 6597 Table 1. Detection Limits in RIU for All Homogeneous Refractive Index Sensing Experiments, Categorized by a Optical Source and Power Level Low Power High Power −7 ECL 5 × 10 9 × 10−8 FPL 1 × 10−6 5 × 10−7 a Each LOD represents the average value of three experiments. compared to the purified water (Milli-Q) flowing buffer [37]. A flow velocity of F = 30 µL/min was maintained throughout the experiments. An example of the induced phase shifts due to the four different sodium chloride concentrations evalu- ated with low power by both light sources is shown in Fig. 2. The blue solid and the orange dashed curves correspond to the ECL and the FPL source, respectively. Both are highly corre- Fig. 3. Steady-state equilibrium phase shifts due to additive CRP lated and produce a similar total phase response. The mean immunoreactions using the FPL. The negligible shift of the control extracted sensitivity is S ∼ 7100 rad/RIU which is close to (BSA with a concentration of 4000 ng/mL) is shown in red. For reasons the theoretical value of S = (2π/λ)SWG L s ≈7800 rad/RIU, of visibility, the two highest CRP injections are not shown. “On” and with SWG the waveguide sensitivity, following the definition “Off” indicate the estimated start and end of the injections. The actual in [34]. The inset shows a 60 s fragment of the read-out noise, injection exposure time is likely longer due to diffusion within the and the corresponding standard deviation (σ ). These exper- fluidic system. iments were carried out three times, averaging the resulting sensitivities and noise standard deviations to compute the LODs shown in Table 1 as LOD = 3σ/S. Considering the highly distinct spectra of the two sources [Fig. 1(a)], it is remarkable that both extracted LODs are on the same order of magni- tude, differing just by a factor of ≈2 − 3 at the same power level. The ECL experiences a LOD improvement by a factor of ≈6 when the input power is switched from low to high, the FPL a factor of ≈2. The fact that they do not produce the same LOD is attributed mainly to the fact that the low-cost FPL is poorly temperature controlled. This can result in a drift of the central wavelength, which, combined with the wavelength response of the grating couplers, can introduce an unwanted modulation. To validate the performance of the device as a biosensor, benchmark CRP protein detection experiments were per- formed with the high power setting. CRP was chosen because it is a widely used biomarker for inflammatory conditions Fig. 4. CRP calibration curve obtained for triplicate immunosens- [38]. Prior to each experiment, the following already success- ing experiments with ECL and FPL, operated at higher power level. fully established offline chemical procedure was carried out The data shown is the mean (filled circle) and standard deviation (error [39]: the sensor chip surface was first oxidized by immersion bars). The lines represent a nonlinear fit to the mean data. in piranha solution, and subsequently amine-functionalized with (3-aminopropyl)trimethoxysilane (APTMS). Finally, anti-CRP antibodies were covalently attached to the surface 200 ng/mL follow a linear trend. However, as CRP accumu- via an amine-to-amine linker, bis(sulfosuccinimidyl)suberate lates on the surface, the biorecognition layer starts to saturate, (BS3). Before each detection experiment, the sensor surface and increasing concentrations yield smaller phase shifts. The was blocked with bovine serum albumin (BSA, 10 mg/mL), corresponding mean calibration curves for CRP detection, to suppress non-specific binding. A flow rate of 20 µL/min for both the low-cost FPL and the reference ECL source, are was established for all CRP additive immunosensing experi- represented in Fig. 4. The data of the cumulative introduced ments with a phosphate-buffered saline (PBS) 10 mM (pH 7.4) phase changes 1ϕcum for each experiment were fitted to the as running buffer. Seven different CRP concentrations were function BC cum /(D + C cum ) achieving an excellent fitting subsequently injected to estimate the LOD: 10, 50, 200, 500, (R 2 > 0.99 for each calibration), with C cum the cumulative 1000, 2000, 4000 ng/mL. Prior to that, BSA was injected at a sum of the injected CRP concentrations, and B, D the fitting concentration of 4000 ng/mL to confirm specificity. Figure 3 parameters. The dots represent the mean cumulative phase shift shows the sensors phase response according to the injected and the error bars the standard deviation for the calibration CRP concentration using the low-cost FPL. The BSA con- curve for the corresponding cumulative concentration. 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