This is the accepted version of the article: Leuermann J., Stamenkovic V., Ramirez-Priego P., Sánchez-Postigo A., Fernández-Gavela A., Chapman C.A., Bailey R.C., Lechuga L.M., Perez-Inestrosa E., Collado D., Halir R., Molina-Fernández Í.. Coherent silicon photonic interferometric biosensor with an inexpensive laser source for sensitive label-free immunoassays. Optics Letters, (2020). 45. : - . 10.1364/OL.411635. Available at: https://dx.doi.org/10.1364/OL.411635 Letter Optics Letters 1 Coherent Silicon Photonic Interferometric Biosensor with Inexpensive Laser Source for Sensitive Label-Free Immunoassays J ONAS L EUERMANN1,2* , V LADIMIR S TAMENKOVIC1,3 , PATRICIA R AMIREZ -P RIEGO4 , A LEJANDRO S ÁNCHEZ -P OSTIGO2 , A DRIÁN F ERNÁNDEZ -G AVELA5 , C OLE A. C HAPMAN6 , RYAN C. BAILEY6 , L AURA M. L ECHUGA4 , E ZEQUIEL P EREZ -I NESTROSA1,3 , DANIEL C OLLADO1,3 , R OBERT H ALIR1,2 , AND Í ÑIGO M OLINA -F ERNÁNDEZ1,2 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, MI 48109, USA * Corresponding author: jonas.leuermann@uma.es Compiled October 30, 2020 Over the last two decades integrated photonic sensors principle of evanescent field sensing: the analyte attaches to a have been of major interest to the optical biosensor biorecognition layer on the waveguide surface, where it interacts community due to their capability to detect low concen- with the evanescent field of the guided wave, changing its phase trations of molecules with label-free operation. Among [6–8]. Since changes in the optical phase cannot be directly mea- these, interferometric sensors can be read-out with sim- sured, they are converted into changes in optical power, using ple, fixed-wavelength laser sources and offer excellent on-chip resonant or interferometric structures. Resonant struc- tures convert the phase change into a resonance wavelength detection limits, but can suffer from sensitivity fad- shift, thus often requiring a read-out system with a tunable ex- ing when not tuned to their quadrature point. Re- ternal cavity laser [9–14], or a broadband light source combined cently, coherently detected sensors were demonstrated with additional wavelength filtering [15]. For homogeneous as an attractive alternative to overcome this limitation. refractive index sensing these systems offer moderate limits of Here we show, for the first time, that this coherent detection (LOD), in the order of 10−6 refractive index units (RIU), scheme provides sub-nanogram per milliliter limits of but can achieve very compact footprints [9, 16, 17]. Regarding detection in C-reactive protein immunoassays, and that the detection of biomolecules, in benchmark primary C-reactive quasi-balanced optical arm lengths enable operation protein (CRP) binding assays resonant sensors achieve LODs with inexpensive Fabry-Perot type lasers sources at tele- in the order of 10 ng/mL (0.4 nM) [18]. Interferometric sensors com wavelengths. © 2020 Optical Society of America compare the optical wave in the sensing waveguide to a wave in a reference waveguide to reconstruct the phase shift. These sen- http://dx.doi.org/10.1364/ao.XX.XXXXXX sors generally operate with a fixed-spectrum laser source, which can be a superluminescent diode, with asymmetric sensing and Integrated photonic biosensors have been the subject of intense reference arm and optical wavelength filtering in the read-out research in the last two decades, not only due to their minituriz- [2, 19, 20], or a narrow-linewidth laser source, e.g. a He-Ne ability but also because they can quantitatively detect extremely laser at visible wavelengths [21–23] or an external cavity lasers low concentrations of analytes in real time and without the at telecom wavelengths [10, 24, 25]. Detection limits down to need of labeling tags [1–3]. Silicon based sensors provide par- 2 ng/mL (84 pM) with label-free primary binding immunoassay ticularly compact footprints, enabling multiplexed detection of for detecting CRP have been reported for Mach-Zehnder interfer- different biochemical substances, and CMOS compatible fabrica- ometers [16, 26]. It is noteworthy that shortcomings commonly tion, which paves the way towards low-cost mass production associated with these interferometric sensors, such as sensitivity [4]. Therefore, they are candidates for point-of-care (POC) solu- fading when not operated at the quadrature point and phase tions in which pre-functionalized, disposable chips are used in ambiguity, have been overcome, either by adding a modulation conjunction with a read-out system to provide diagnostics that signal [27–29] or by using a coherent detection approach, illus- would otherwise require specialized laboratories and trained trated in Fig. 1a [24, 30–34]. With coherent detection, no tuning personnel [5]. Most photonic integrated sensors rely on the to a quadrature point is required, so that the optical lengths of Letter Optics Letters 2 a 𝜋 External Cavity 12% ∝ exp 𝑗Δ𝜑 Fabry-Perot Δ𝜑 rad Fiber Sensing Arm 0 9% Grating Reference Arm −𝜋 𝜆0 𝜆 b 6% Fabry-Perot Laser 𝜎 ≈ 1.2 mrad External Cavity Laser 𝑃in 𝑃in 𝛿𝜆 𝛿𝜆 Δ𝜆 𝜎 ≈ 3.2 mrad 3% 𝜆0 𝜆 𝜆0 𝜆 Fig. 1. a Schematic of a coherently read balanced Mach- Zehnder sensor, where the phase difference between the ref- erence and the sensing arm is recovered from three output signals. The wavelength response of such a sensors is com- Fig. 2. Bulk sensing using the reference ECL source (continu- pletely flat when the optical arm lengths are balanced. b The ous blue curve) and the low-cost FPL (dashed orange curve) approximated laser spectra of a narrow linewidth external cav- for 3%, 6%, 9%, and 12% mass percentage sodium chloride ity laser and a Fabry-Perot laser, with center wavelength λ0 , injections, respectively, and low input power. linewidth δλ, and spectral spacing ∆λ. our read-out system the light source is connected to an optional the interferometer arms can be perfectly matched. So far, this erbium doped fiber amplifier and a variable optical attenuator approach has only been demonstrated with bulk sensing exper- to obtain the desired input power. A polarization rotator is used iments [10, 24], and using high quality external cavity lasers to achieve TE polarization that is coupled into the silicon nitride with narrow linewidths (see Fig. 1b). Here we show, for the sensor which contains a grating coupler, with coupling losses of first time, that this coherent approach is well suited for protein ∼ 10 dB (see Fig. 1a). The sensor chip is fabricated with a multi- detection, reporting a limit of detection better than 300 pg/mL project wafer approach in a 300 nm thick silicon nitride film for primary binding CRP detection. We furthermore show that [36]. The sensing arm is covered with a polydimethylsiloxane due to quasi-identical optical arm lengths, the sensor can operate (PDMS) microfluidic channel, with a width of 1 mm, a length of with a low-cost Fabry-Perot telecom laser (see Fig. 1b) with an 9.8 mm, and a height of 0.5 mm, which is connected to a syringe acceptable performance penalty. Mach-Zehnder based sensors pump. In between, a six-port valve system enables the injection split the incoming light into two single-mode waveguides, one of an analyte volume of Vinj ≈ 120 µL into the channel. The acting as the sensing arm and the other acting as the reference three optical output signals are coupled to a fiber array using arm (see Fig. 1a). The reference waveguide is buried in silicon on-chip identical grating couplers and are detected by separate dioxide, so that the propagating mode always experiences a photodiodes (PD) with integrated transimpedance amplifiers. constant phase shift. In the sensing waveguides the propagat- The resulting voltages are quantized by a data acquisition board, ing mode interacts with the analyte which causes an additional and signal processing is used to estimate the phase change ∆ϕ phase shift. Conventional Mach-Zehnder interferometers recom- [31]. A detailed description and bulk LOD optimization of the bine these two arms into one output waveguide (e.g. using a experimental setup can be found in [24]. To assess the impact Y-Junction or a 1 × 2 multi-mode interference coupler). This re- of the laser linewidth on the sensor performance, two different sults in a signal proportional to sin (∆ϕ), where ∆ϕ is the phase lights sources were employed in the experiments. As a high difference between the waves in the two arms. This signal is quality reference, we used the WSL-100 from SANTEC (Japan), affected by relative intensity noise, and causes sensitivity fading a single mode, external cavity laser, bought for approximately and phase ambiguity [16]. In contrast, a coherently read MZI 10, 000€, with a linewidth of δλ < 80 fm (100 kHz) (see Fig. 1b). recombines the sensing and reference arms via a 2 × 3 coupler As a low cost alternative we used a hand-held Fabry-Perot laser (see Fig. 1a). From the three output powers a complex signal with a center wavelength of λ0 = 1.543 µm, ∆λ = 1.36 nm wave- ∝ exp (j∆ϕ) can be calculated [31, 35], with the extracted phase length spacing, and a bandwidth of δλ ∼ 3.6 nm (450 GHz) (see being immune to relative intensity noise, sensitivity fading and Fig. 1b), purchased on Amazon for 34€. In the following we phase ambiguity [34].The noise in the extracted p phase caused by will refer to these sources as ECL (External Cavity Laser) and a finite laser linewidth δ f is given by 2∆τ π δ f B, where B is FPL (Fabry-Perot Laser), respectively. Two types of experiments the measurement bandwidth and ∆τ is the time delay between were chosen to determine the quality of the sensor, bulk refrac- the reference and sensing arm, ∆τ = | Ls ng,s − Lr ng,r |/c0 with tive index sensing and protein CRP immunodetection. We used Ls , Lr the physical lengths of the sensing and reference arms, two different power levels: low, with a total optical power of ng,s , ng,r the group indices, and c0 is the speed of light [34]. Thus, approximately of −30 dBm received by the PDs and high, with by choosing Ls and Lr such that ∆τ = 0, the phase noise can a power of approximately −10 dBm. Homogeneous refractive be minimized. Intuitively, this can be understood because for index sensing experiments were performed by injecting four a perfectly balanced interferometer the wavelength response is different mass percentage solutions of sodium chloride (3%, 6%, flat (see Fig. 1a). In our experiments the reference and sens- 9%, 12%) resulting in changes of (0.5, 1.0, 1.5, 2.0) × 10−2 RIU ing arm have the same length L = 8.8 mm and a group index compared to the purified water (Milli-Q) flowing buffer [37]. A difference of ∆ng,rs = |ng,s − ng,r | ≈ 0.01 at λ0 = 1.55 µm. In flow velocity of F = 30 µL/min was maintained throughout Letter Optics Letters 3 2500 2000 “On” “Off” 1500 1000 500 0 0 2000 4000 6000 8000 Fig. 3. Steady-state equilibrium phase shifts due to additive Fig. 4. CRP calibration curve obtained for triplicate im- CRP immunoreactions using the FPL. The negligible shift munosensing experiments with ECL and FPL, operated at of the control (BSA with a concentration of 4000 ng/mL) is higher power level. The data in purple shows the mean (filled shown in red. For reasons of visibility, the two highest CRP circle) and the standard deviation (errorbars). The lines repre- injections are not shown. "On" and "Off" indicate the estimated sent a nonlinear fit to the mean data. start and end of the injections. The actual injection exposure time is likely longer due to diffusion within the fluidic system. were performed with the high power setting. CRP was chosen the experiments. An example of the induced phase shifts due because it is a widely used biomarker for inflammatory con- to the four different sodium chloride concentrations evaluated ditions [38]. Prior to each experiment, the following already with low power by both light sources is shown in Fig. 2. The successfully established offline chemical procedure was carried blue solid and the orange dashed curves correspond to the ECL out [39]: the sensor chip surface was first oxidized by immer- and the FPL source, respectively. Both are highly correlated sion in piranha solution, and subsequently amine-functionalized and produce a similar total phase response. The mean extracted with (3-aminopropyl)trimethoxysilane (APTMS). Finally, anti- sensitivity is S ∼ 7100 rad/RIU which is close to the theoret- CRP antibodies were covalently attached to the surface via an ical value of S = (2π/λ)SWG Ls ≈ 7800 rad/RIU, with SWG amine-to-amine linker, bis(sulfosuccinimidyl)suberate (BS3). Be- the waveguide sensitivity, following the definition in [34]. The fore each detection experiment, the sensor surface was blocked inset shows a 60 seconds fragment of the read-out noise, and with bovine serum albumin (BSA, 10 mg/mL), to suppress non- the corresponding standard deviation (σ). These experiments specific binding. A flow rate of 20 µL/min was established for were carried out three times, averaging the resulting sensitiv- all CRP additive immunosensing experiments with a phosphate- ities and noise standard deviations to compute the limits of buffered saline (PBS) 10 mM (pH 7.4) as running buffer. Seven detection shown in Table 1 as LOD = 3σ/S. Considering different CRP concentrations were subsequently injected to es- timate the LOD: 10, 50, 200, 500, 1000, 2000, 4000 ng/mL. Prior to that, BSA was injected at a concentration of 4000 ng/mL to Table 1. Detection limits in RIU for all homogeneous re- confirm specificity. Figure 3 shows the sensors phase response fractive index sensing experiments, categorized by optical according to the injected CRP concentration using the low-cost source and power level. Each LOD represents the average FPL. The BSA control injection causes a negligible phase shift, value of three experiments. indicating the CRP specificity of the immunosensor surface. The phase shifts due to the CRP injections with concentrations of Low Power High Power 10, 50, and 200 ng/mL follow a linear trend. However, as CRP ECL 5 × 10−7 9 × 10−8 accumulates on the surface, the biorecognition layer starts to saturate and increasing concentrations yield smaller phase shifts. FPL 1 × 10−6 5 × 10−7 The corresponding mean calibration curves for CRP detection, for both the low-cost FPL and the reference ECL source, are rep- the highly distinct spectra of the two sources (see Fig. 1a), it resented in Fig. 4. The data of the cumulative introduced phase is remarkable that both extracted LODs are in the same order changes ∆ϕcum for each experiment was fitted to the function of magnitude, just differing by a factor of ≈ 2 − 3 at the same BCcum /( D + Ccum ) achieving an excellent fitting (R2 > 0.99 power level. The ECL experiences a LOD improvement by a for each calibration), with Ccum the cumulative sum of the in- factor of ≈ 6 when the input power is switched from low to jected CRP concentrations, and B, D the fitting parameters. The high, the FPL a factor of ≈ 2. The fact that they do not produce dots represent the mean cumulative phase shift and the error- the same LOD is mainly attributed to the fact that the low-cost bars the standard deviation for the calibration curve for the FPL is poorly temperature controlled. This can result in a drift of corresponding cumulative concentration. The high standard the central wavelength, which, combined with the wavelength deviation of the sensitivity can be explained by two factors. response of the grating couplers can introduce an unwanted First, for each experiment a different chip was used, each of modulation. In order to validate the performance of the device which chip had been functionalized several times before, pos- as a biosensor, benchmark CRP protein detection experiments sibly degenerating the antibody immobilization efficiency. Sec- Letter Optics Letters 4 ond, but more importantly, our functionalization strategy re- 12. S. Hu, Y. Zhao, K. Qin, S. T. Retterer, I. I. Kravchenko, and S. M. Weiss, quires further optimization to achieve reliable homogeneous ACS Photonics 1, 590 (2014). surface coverage. A worst case sensitivity of 2 mrad/(ng/mL) 13. Q. Liu, X. Tu, K. W. Kim, J. S. Kee, Y. Shin, K. Han, Y.-J. Yoon, G.-Q. and detection limit just below 300 pg/mL was extracted from Lo, and M. K. Park, Sens. 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Bioelectron. 22, 2591 (2007). We acknowledge funding from the European Union’s Hori- 21. S. Dante, Opt. Pura y Apl. 45, 87 (2012). zon 2020 Marie Skłodowska-Curie grant agreement No 713721; 22. K. Zinoviev, L. G. Carrascosa, J. S. del Río, B. Sepúlveda, the Spanish Ministry for Science, Innovation, and Universi- C. Domínguez, and L. M. Lechuga, Adv. Opt. Technol. 2008, 1 (2008). ties (cofinanciado FEDER), TEC2016-80718-R, CTQ2016-75870-P, 23. N. Skivesen, A. Têtu, M. Kristensen, J. Kjems, L. H. Frandsen, and P. I. RD16/0006/0012 and PID2019-104293GB-I00. Proyecto I+D+i Borel, Opt. Express 15, 3169 (2007). en el marco del Programa Operativo FEDER Andalucía 2014- 24. J. Leuermann, A. Fernández-Gavela, A. Torres-Cubillo, S. Postigo, A. Sánchez-Postigo, L. M. Lechuga, R. Halir, and Í. Molina-Fernández, 2020 (UMA18-FEDERJA-219 and UMA18-FEDERJA-007), Junta Sensors 19, 3671 (2019). de Andalucía, Proyectos Excelencia-Retos (P18-RT-1453, P18-RT- 25. A. Densmore, D.-X. Xu, S. Janz, P. Waldron, T. Mischki, G. Lopinski, 793); Severo Ochoa program from Spanish MINECO (Grant No. A. Delâge, J. Lapointe, P. Cheben, B. Lamontagne, and J. H. Schmid, SEV-2017-0706). CTQ2017-86994-R project from the Spanish Min- Opt. Lett. 33, 596 (2008). istry of Economy and Competitiveness (MINECO). European 26. A. Psarouli, A. Botsialas, A. Salapatas, G. Stefanitsis, D. Nikita, Union’s Horizon 2020 research and innovation programme un- G. Jobst, N. Chaniotakis, D. Goustouridis, E. Makarona, P. S. Petrou, der EuroNanoMed 3-H2020 DrNanoDAII and MINECO, PRO- I. Raptis, K. Misiakos, and S. E. Kakabakos, Talanta. 165, 458 (2017). GRAMACION CONJUNTA 2019/PCI 2019-2. 27. S. Dante, D. Duval, B. Sepúlveda, A. B. González-Guerrero, J. R. Sendra, and L. M. Lechuga, Opt. Express 20, 7195 (2012). Disclosures. The authors declare no conflicts of interest. 28. A. B. González-Guerrero, J. Maldonado, S. Herranz, and L. M. Lechuga, Anal. Methods 8, 8380 (2016). 29. Y. E. Marin, V. Toccafondo, P. Velha, Y. 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