Integration of microfluidics with grating coupled silicon photonic sensors by one-step combined photopatterning
and molding of OSTE
Carlos Errando-Herranz,1,2Farizah Saharil,1Albert Mola Romero,1,3 Niklas Sandstr¨om,1Reza Zandi Shafagh,1Wouter van der Wijngaart,1
Tommy Haraldsson,1and Kristinn B. Gylfason1,∗
1KTH Royal Institute of Technology, Micro and Nanosystems, Stockholm, Sweden
2UPV Polytechnic University of Valencia, Valencia, Spain
3University of Barcelona, Barcelona, Spain
* kristinn.gylfason@ee.kth.se
Abstract: We present a novel integration method for packaging silicon photonic sensors with polymer microfluidics, designed to be suitable for wafer-level production methods. The method addresses the previously unmet manufacturing challenges of matching the microfluidic footprint area to that of the photonics, and of robust bonding of microfluidic layers to biofunctionalized surfaces. We demonstrate the fabrication, in a single step, of a microfluidic layer in the recently introduced OSTE polymer, and the subsequent unassisted dry bonding of the microfluidic layer to a grating coupled silicon photonic ring resonator sensor chip. The microfluidic layer features photopatterned through holes (vias) for optical fiber probing and fluid connections, as well as molded microchannels and tube connectors, and is manufactured and subsequently bonded to a silicon sensor chip in less than 10 minutes. Combining this new microfluidic packaging method with photonic waveguide surface gratings for light coupling allows matching the size scale of microfluidics to that of current silicon photonic biosensors. To demonstrate the new method, we performed successful refractive index measurements of liquid ethanol and methanol samples, using the fabricated device. The minimum required sample volume for refractive index measurement is below one nanoliter.
© 2013 Optical Society of America
OCIS codes: (130.6010) Sensors; (130.6622) Subsystem integration and techniques;
(160.5470) Polymers.
References and links
1. Y. S. Sun, J. P. Landry, Y. Y. Fei, X. D. Zhu, J. T. Luo, X. B. Wang, and K. S. Lam, “Effect of fluorescently labeling protein probes on kinetics of protein-ligand reactions.” Langmuir 24, 13399–13405 (2008).
2. T. Kodadek, “Protein microarrays: prospects and problems,” Chemistry & Biology 8, 105–115 (2001).
3. G. Zheng, F. Patolsky, Y. Cui, W. U. Wang, and C. M. Lieber, “Multiplexed electrical detection of cancer markers with nanowire sensor arrays,” Nat. Biotechnol. 23, 1294–1301 (2005).
4. J. M. Rothberg, W. Hinz, T. M. Rearick, J. Schultz, W. Mileski, M. Davey, J. H. Leamon, K. Johnson, M. J.
Milgrew, M. Edwards, J. Hoon, J. F. Simons, D. Marran, J. W. Myers, J. F. Davidson, A. Branting, J. R. Nobile, B. P. Puc, D. Light, T. A. Clark, M. Huber, J. T. Branciforte, I. B. Stoner, S. E. Cawley, M. Lyons, Y. Fu, N. Homer,
M. Sedova, X. Miao, B. Reed, J. Sabina, E. Feierstein, M. Schorn, M. Alanjary, E. Dimalanta, D. Dressman, R. Kasinskas, T. Sokolsky, J. A. Fidanza, E. Namsaraev, K. J. McKernan, A. Williams, G. T. Roth, and J. Bustillo,
“An integrated semiconductor device enabling non-optical genome sequencing,” Nature 475, 348–352 (2011).
5. K. De Vos, I. Bartolozzi, E. Schacht, P. Bienstman, and R. Baets, “Silicon-on-insulator microring resonator for sensitive and label-free biosensing,” Opt. Express 15, 7610–7615 (2007).
6. M. Iqbal, M. A. Gleeson, B. Spaugh, F. Tybor, W. G. Gunn, M. Hochberg, T. Baehr-Jones, R. C. Bailey, and L. C.
Gunn, “Label-free biosensor arrays based on Silicon ring resonators and high-speed optical scanning instrumen- tation,” IEEE J. Quantum Electron. 16, 654–661 (2010).
7. A. Densmore, D. X. Xu, S. Janz, P. Waldron, T. Mischki, G. Lopinski, A. Delˆage, J. Lapointe, P. Cheben, B. Lam- ontagne, and J. H. Schmid, “Spiral-path high-sensitivity silicon photonic wire molecular sensor with temperature- independent response,” Opt. Lett. 33, 596–598 (2008).
8. O. Parriaux, V. A. Sychugov, and A. V. Tishchenko, “Coupling gratings as waveguide functional elements,” Pure and Appl. Opt.: J. European Optical Society Part A 5, 453+ (1999).
9. M. Antelius, K. B. Gylfason, and H. Sohlstr¨om, “An apodized SOI waveguide-to-fiber surface grating coupler for single lithography silicon photonics,” Opt. Express 19, 3592–3598 (2011).
10. M. S. Luchansky and R. C. Bailey, “High-Q optical sensors for chemical and biological analysis,” Anal. Chem.
84, 793–821 (2011).
11. K. De Vos, J. Girones, T. Claes, Y. De Koninck, S. Popelka, E. Schacht, R. Baets, and P. Bienstman, “Multiplexed antibody detection with an array of Silicon-on-insulator microring resonators,” IEEE Photonics J. 1, 225–235 (2009).
12. C. F. Carlborg, K. B. Gylfason, A. Kazmierczak, F. Dortu, M. J. Banuls Polo, A. Maquieira Catala, G. M.
Kresbach, H. Sohlstrom, T. Moh, L. Vivien, J. Popplewell, G. Ronan, C. A. Barrios, G. Stemme, and W. van der Wijngaart, “A packaged optical slot-waveguide ring resonator sensor array for multiplex label-free assays in labs-on-chips,” Lab on a Chip 10, 281–290 (2010).
13. C. F. Carlborg, T. Haraldsson, K. Oberg, M. Malkoch, and W. van der Wijngaart, “Beyond PDMS: off- stoichiometry thiol-ene (OSTE) based soft lithography for rapid prototyping of microfluidic devices,” Lab on a Chip 11, 3136–3147 (2011).
14. F. Saharil, C. F. Carlborg, T. Haraldsson, and W. van der Wijngaart, “Biocompatible ”click” wafer bonding for microfluidic devices,” Lab on a Chip 12, 3032–3035 (2012).
15. J. M. Karlsson, F. Carlborg, F. Saharil, F. Forsberg, F. Niklaus, W. van der Wijngaart, and T. Haraldsson, “High- resolution micropatterning of off-stoichiometric thiol-enes (OSTE) via a novel lithography mechanism,” in “16th International Conference on Miniaturized Systems for Chemistry and Life Sciences (MicroTAS),” (2012).
16. B. Movassagh and M. Soleiman-Beigi, “Synthesis of thiocarbamates from thiols and isocyanates under catalyst- and solvent-free conditions,” Monatshefte f¨ur Chemie 139, 137–140 (2008).
17. C. F. Carlborg, M. Cretich, T. Haraldsson, L. Sola, M. Bagnati, M. Chiari, and W. van der Wijngaart, “Biosticker:
patterned microfluidic stickers for rapid integration with microarrays,” in “15th International Conference on Miniaturized Systems for Chemistry and Life Sciences (MicroTAS),” (2011), pp. 311–313.
18. N. Sandstrom, R. Z. Shafagh, C. F. Carlborg, T. Haraldsson, G. Stemme, and W. van der Wijngaart, “One step integration of gold coated sensors with OSTE polymer cartridges by low temperature dry bonding,” in “16th International Conference on Solid-State Sensors, Actuators and Microsystems,” (IEEE, 2011), pp. 2778–2781.
19. C. Errando-Herranz, F. Saharil, A. Mola Romero, N. Sandstr¨om, R. Z. Shafagh, W. van der Wijngaart, T. Har- aldsson, and K. B. Gylfason, “Integration of polymer microfluidic channels, vias, and connectors with Silicon photonic sensors by one-step combined photopatterning and molding of OSTE,” in “17th International Confer- ence on Solid-State Sensors, Actuators and Microsystems,” (IEEE, 2013), pp. 1613–1616.
20. E. P. Kartalov and S. R. Quake, “Microfluidic device reads up to four consecutive base pairs in DNA sequencing- by-synthesis,” Nucleic Acids Research 32, 2873–2879 (2004).
21. C. F. Carlborg, T. Haraldsson, M. Cornaglia, G. Stemme, and W. van der Wijngaart, “A High-Yield Process for 3-D Large-Scale Integrated Microfluidic Networks in PDMS,” J. Microelectromech. Syst. 19, 1050–1057 (2010).
22. D. R. Lide, ed., CRC Handbook of Chemistry and Physics (CRC, 2008).
23. J. E. Bertie and Z. Lan, “Infrared intensities of liquids XX: the intensity of the OH stretching band of liquid water revisited, and the best current values of the optical constants of H2O(l) at 25C between 15,000 and 1 cm-1,” Appl.
Spectrosc. 50, 1047–1057 (1996).
1. Introduction
The combination of a biological recognition element with a physical transducer makes biosens- ing a powerful tool for biological and medical analysis. Although widely used, label-based biosensing has several limitations, such as the risk of interference with the reaction under study [1], labeling heterogeneity [2], and a lack of real-time kinetics measurements for quan- tifying reaction rates. Label-free biosensing addresses these limitations by providing real-time
physically quantifiable information without labels.
Silicon based sensors have recently shown utility for important biological and chemical label- free sensing applications. Silicon nanowire sensors [3], and transistor based next generation DNA sequencing devices [4], are examples of such sensors showing low detection limits, high scalability, simple fabrication, and mass production capability in dense arrays at low cost.
Silicon photonic waveguide based biosensors share these characteristics. These sensors de- tect refractive index changes within their evanescent field upon biomolecule binding with high sensitivity, and can be fabricated in compact arrays using standard lithography techniques. Sev- eral such biosensors have been reported: Ring resonators [5] have been shown to be scalable to large sensing arrays [6], and yield volume refractive index sensitivities up to 70 nm/RIU [5].
Spiral-path interferometers have shown sensitivity of 163 nm/RIU [7]. Moreover, the use of sur- face grating couplers enables the probing of light everywhere on chip [8]. Grating couplers can be added to the sensing circuits without an additional lithography step, yielding a one-step fab- rication process for grating coupled photonic sensor chips on SOI substrates [9]. These reports show that silicon photonic label-free biosensing has the potential to become a highly scalable and low cost sensing technique. However, for chemical and biological sensing, the integration of liquid handling systems consumes a considerably larger wafer area than that of the photonic footprint, and thus the scaling benefit is lost.
Molded polydimethylsiloxane (PDMS) constitutes the current academic solution for sili- con photonic biosensor microfluidics [10], but drawbacks such as large wafer area consump- tion, long curing times, adsorption of small biomolecules into the PDMS, and lack of bonding techniques compatible with surface biofunctionalization, make industrial application question- able [10]. These limitations are apparent in [11], in which stamping with an epoxy glue for bonding results in channel clogging. Moreover, PDMS based soft-lithography molding of vias (through holes) is hampered by squeeze-film formation, thus necessitating a second low reso- lution via fabrication step by hole punching [12].
To address these limitations, the Off-Stoichiometry Thiol-Ene (OSTE) polymer was intro- duced [13] and microfluidic integration on silicon at wafer level demonstrated [14]. More recently, by a UV-initiated thiol-ene reaction, vias have been photopatterned within a few seconds [15], and a low temperature dry-bonding technique enabled by a room-temperature thiol-isocyanate click reaction [16], permitting bonding to biofunctionalized silicon [17] and gold [18]. Moreover, we showed integration of OSTE onto a silicon photonic interferometer sensor, by coupling light to the chip and verifying the interference in the output power [19].
Here, we demonstrate the integration of OSTE on silicon photonics by performing refrac- tive index measurements with a silicon photonic ring resonator chip. We present a one-step integration of microfluidics onto silicon photonic sensors able to match the size scale of the liquid handling system with that of the silicon photonics, by using a self-bonding OSTE poly- mer microfluidic layer that is structured using a combined photolithography and micromolding process. Figure 1 illustrates the process: 1) By a photolithography of OSTE, we open optical vias with the same size scale as optical fibers above grating couplers, enabling light coupling anywhere on the sensing chip. 2) For rapid prototyping, we mold tube connectors directly into the microfluidic layer. These tube connectors can be substituted by compact manifold connec- tors, if needed, for a higher integration density at wafer level. 3) We dry bond the microfluidic layer to patterned silicon surfaces by click chemistry.
2. Fabrication
Figure 2 shows a schematic cross section of the microfluidic layer fabrication and bonding to the silicon chip. We used a 70% thiol excess OSTE polymer (OSTE-70) [13] for the microfluidic layer. Along with the general OSTE properties discussed above, OSTE-70 provides an excess
OSTE microfluidic layer
Silicon photonic chip Glass mask
fabrication
PDMS mold fabrication
Dry bonding
Integrated chip ready for photonic sensing 1 step Si
lithography process TOP mold +
photomask
BOTTOM mold
Molding and photolithography of OSTE
500µm
chromium pattern for fluidic via
chromium pattern for optical via SU8 relief 500µm
fluid sample input
fluid sample output light light out
in
Fig. 1. The integration scheme: The top mold is a glass mask combining chromium patterns for photolithography of vias with SU8 reliefs for microchannel molding. The insets show photographs of the reliefs and via patterns on the mask. The bottom PDMS mold defines the chip outline and the fluidic connectors. After UV curing, development, and dry-bonding of the microfluidic layer to the photonic chip, the integrated chip is ready for measurements.
PDMS mold PDMS mold
UV curing
Silicon chip tube
connector
optical fiber optical
via
fluid via
Chromium dot OSTE
prepolymer glass
mask
+
R SH R’
thiol -ene
70˚C
A B
C
Silicon sensor Surface grating coupler
SU8 relief
UV-Initiator R S H
R’
70% thiol R SH excess
SHSH SH SH NCO NCO NCO NCO
S NCO
S NCO
S NCO
S NCO
thiol excess
isocyanate
covalent bonding process D
flip over
OSTE
OSTE
Curing 13 s
Developing 30 s Bonding 10 min
Fig. 2. A schematic cross-section of the fabrication of the OSTE microfluidic layer and its bonding to a photonic silicon chip: (a) OSTE-70 is poured into a PDMS mold. (b) A UV cure through a glass mask polymerizes exposed parts by a thiol-ene reaction. (c) The fluidic and optical vias are developed in butyl acetate. (d) A silanized silicon photonic chip is then aligned and bonded to the fluidic layer and cured for 10 min at 70◦C.
of thiol that permits a strong bond to a silanized silicon surface [14].
The OSTE-70 was mixed in a ratio of functional groups of 1.7:1, pentaerythri- tol tetrakis (2-mercaptoacetate) and triallyl-1,3,5-triazine-2,4,6(1H,3H,5H)-trione respec- tively (Sigma-Aldrich). We added 0.1% (by mass) of the photoinitiator ethyl-2,4,6- trimethylbenzoylphenylphosphinate (BASF AG, Germany).
We poured the OSTE-70 prepolymer into a PDMS mold defining the chip outline and fluid connectors (Fig. 2(a)), and then sandwiched the polymer between the PDMS and a glass mask.
The glass mask contains chromium patterns that define vias, and 50 µm thick SU8 reliefs that define microchannels, and thus acts simultaneously as a mold and a photolithography mask (Fig. 2(b)). The chromium patterns for fluidic connections on the glass mask were aligned by
tube connector
optical via microchannel
optical fibers rubber tubing
molded microchannel photolithographed optical via molded tube connector
2 mm
photolithographed fluid via A
microchannelOSTE
bondarea microchannel surface grating
surface grating
120 µm Ring resonator
optical via
B
grating taper
ring resonator 500 nm
microchannel bond area 75 Pm
C
2 mm
75 µm
Fig. 3. (a) A photograph of the ring resonator chip. In the magnified image, the surface grat- ing couplers of a ring resonator sensor device are visible. (b) The design of one of the ring resonator sensors fabricated in the silicon chip and the OSTE microchannel dimensions.
(c) The measurement setup, with connected optical and fluidic ports.
eye to the connector molds in the PDMS mold. After a 13 s UV exposure at 13 mW/cm2, using a collimated NUV lightsource with wavelength peaks at 365, 405, and 436 nm (OAI, USA), the OSTE layer was released and subsequently developed in butyl acetate for 30 s (Fig. 2(c)).
During development, unexposed prepolymer dissolves, leaving open vias for optical and fluidic connections, as seen in the photographs in Fig. 3.
We functionalized the surface of the silicon photonic chip by dipping it for 10 min into 5% (by mass) of 3-(triethoxysilyl)propyl isocyanate (Sigma-Aldrich) in toluene. We dried the chips for 10 min at 70◦C, and thereafter aligned the microfluidic layer to the chips using a long working distance microscope objective and a micropositioning stage, followed by bond- ing. The bonding takes advantage of the excess thiol functional groups on the OSTE surface that covalently bond the fluidic layer to the silicon chip, by a click reaction [16] during 10 min at 70◦C (Fig. 2(d)). We have already reported a reduction of the bond time down to 5 min and temperature down to 37◦C using this method [14]. As illustrated in Figs. 3(a) and 3(b), the optical vias are separated by only 75 µm wide bond areas from the 100 µm wide microfluidic channel. This tight spacing allows grating coupler separation of only 500 µm, thus reducing the required wafer footprint of the silicon photonics. Due to the limited accuracy of punch- ing, punched vias in PDMS are commonly separated by close to 1 mm [20]. PDMS via to channel spacing down to 200 µm has been recently shown using a polymerization inhibition technique [21], but this method leaves the PDMS surface unpolymerized, and has only been used to perforate thin membranes and not combined with connector molding, as shown here.
A particular feature of the OSTE-70 is its low glass transition temperature of 37◦C [14]. This allows us to safely unbond the OSTE layer by raising the temperature to 50◦C and pealing the layer off the silicon surface. This permits reuse of the silicon chips. For permanent bonding, an OSTE formulation with a higher glass transition temperature can be used [13].
3. Optical measurements and results
To demonstrate the usefulness of the integrated chip for refractive index based sensing, we measured the resonance wavelength shifts of a silicon photonic ring resonator sensor (40 µm diameter [9]), upon the injection of a dilution series of ethanol and methanol in water. The measurement setup is shown in Fig. 3(c). One fluidic port is connected to a syringe pump in suction mode and samples are introduced by pipetting into the other connector. With 100 µm
1552 1552.5 1553 1553.5 1554 1554.5 1555 1555.5 1556
−52
−50
−48
−46
−44
−42
−40
−38
−36
−34
−32
Wavelength [nm]
Transmitted power [dBm] waterethanol 10%
ethanol 20%
ethanol 40%
1.33 1.335 1.34 1.345 1.35 1.355 1.36 1.365
−0.2 0 0.2 0.4 0.6 0.8 1 1.2 1.4
Resonance wavelength shift [nm]
Refractive index
water ethanol methanol
A B
Fig. 4. (a) Resonance spectra for different concentrations of ethanol in water. The red shift increases as the ethanol concentration increases. (b) The measured resonance wavelength shift shows a linear dependence on the refractive index of the injected solution.
wide and 50 µm high microchannels above the 40 µm diameter ring, the minimum required sample volume for refractive index measurement is below one nanoliter. After alignment of the optical fibers with the grating couplers, light emitted from a laser source is coupled into the chip, and thus the resonance peaks can be observed with a wavelength domain component analyzer (Agilent Technologies 86082A).
We mixed ethanol and methanol in DI water at different concentrations, yielding six solu- tions with a range of refractive indexes from 1.333 to 1.358 RIU [22]. We flowed the samples through the chip at a flow rate of 3.5 µL/min, with a DI water flush between each of them.
The resonance shifts observed for the ethanol solutions are shown in Fig. 4(a). We observe a red shift from 1552.8 nm to 1554.1 nm, and an increase in Q from 17 200 to 17 300, as the ethanol concentration increases. The increase in Q is caused by the reduced infrared absorption of the water in the sample at higher ethanol concentration [23]. Figure 4(b) shows the measured resonance wavelength shift for all the samples, as a function of refractive index. The slopes of the linear fits for ethanol and methanol coincide, yielding a volume refractive index sensitivity of 50.5 nm/RIU, in good agreement with previously reported similar devices [5]. Although the slopes for both solutes are consistent, we observe a slight offset, most likely due to a difference in temperature around 10◦C between the two solution series.
4. Conclusions
We have presented a novel one-step microfluidic integration technique for silicon photonic waveguide based sensors with surface grating couplers. Using the lithographic capability of the OSTE polymer, we combine lithography and molding to enable the integration of pho- tolithographed vias for optical probing and microfluidic channels in a single step. Moreover, dry bonding of OSTE to silanized silicon photonic chips is compatible with biofunctionalized surfaces. We show leakage free bonding to patterned silicon, together with refractive index measurements using a grating coupled ring resonator sensor chip, fabricated by these means.
By using this process, extendable to the wafer scale, we can match the size scale of the OSTE based microfluidics to the current size scale of label-free photonic biosensors.
Acknowledgments
This work was partially supported by the Swedish Research Council (B0460801), the G¨oran Gustafsson Foundation, and the European Research Council (267528).
