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Abstract
In this work we demonstrate the use of a dielectric barrier discharge plasma for the treatment of SU-8. The resulting hydrophilic surface displays a 5° contact angle and (0.40 ± 0.012) nm roughness. Using this technique we also present a proof of concept of IgG and prostate specific antigen biodetection on a thin layer of SU-8 over gold via surface plasmon resonance detection.
© 2018 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
Corrections
4 May 2018: Typographical corrections were made to the body text, Refs. 26–33, and the acknowledgments section.
1. Introduction
The SU-8 is an epoxy that has been widely used in the fabrication of optical telecommunication and biosensor devices [1–3]. The advantage of the implementation of this material is based on attributes such as high refractive index, mechanical and thermal stability, reliability, low loss and easy fabrication. All these advantages make it a desirable material for the implementation of highly sensitive sensing devices due to the possibility of low cost mass production [4–8].
Recently, we demonstrated a new fabrication platform for optical devices based on polymer technology in the visible spectrum through Direct Laser Writer (DLW) system, which presents a highly reliable, rapid prototyping and very low cost fabrication method [9]. To photosensitize SU-8 in the visible spectrum, the commercial photoinitiator H-nu 470 was required, which allows to change the maximum pre-polymerized absorption peak of the polymer from 365 nm to 470 nm wavelength. This modification enables the use of H-line lithography and improves the mechanical and thermal properties of the fabricated structures, providing a more robust and resistant device.
In spite of the additional benefits provided by the fabrication method described above, it is indispensable to take into account that most body fluids have around 90% of water content, which are responsible for the transport of hydrophilic proteins which will eventually be immobilized [10–12] in biosensor applications. The naturally hydrophobic surface of SU-8 will minimize the adsorption of any water soluble components, therefore different methods have been implemented to hydrophilize its surface, specially when microfluidic systems are involved. One of the most widely used techniques employs argon and oxygen plasma [13–17]. Such treatment has shown low contact angle of water droplets (high wetability) and good stability after at least 8 minutes of treatment [13–15]. However, the exposure of nanometer-scale thin surfaces or devices to this type of plasma in vacuum for long periods endangers the integrity of the exposed surface, possibly leading to unrepairable damage [13, 18–22].
Biofunctionalization of SU-8 in bulk and devices such as cantilevers has been demonstrated by the deposition of gold nanoparticles onto the device surface and performing the entire process of biofunctionalization on the metal, avoiding the direct interaction of the analytes with the polymer [23–25]. In this work we characterize the performance of SU-8 + H-nu 470 photoinitiator as an optical biosensor, presenting a reliable, complete and robust platform for fabrication and biofunctionalization of the SU-8 surface. An unconventional method of hydrophilization of the polymer surface based on Dielectric Barrier Discharge (DBD) plasma treatment [26] is demonstrated, which allows the reduction of the contact angle of liquids on the treated surface to 5° after exposure, maintaining its roughness around (0.40 ± 0.012 nm. We also present a proof of concept of Immunoglobulin G (IgG) and Recombinant Human) Prostate Specific Antigen (PSA) biodetection in a few-nanometer layer of SU-8 + H-nu 470 photoinitiator using a Surface Plasmon Resonance (SPR) detection system.
Several works based on polymeric technology for biosensing applications, do not show the biofunctionalization process directly on the polymer surface, in some cases the biofunctionalization is made on materials such as metal (nano-particles or nano-layer) preferably gold or silver as it was said in our manuscript, deposited on polymeric surface [23–25], in others cases, was used titanium dioxide (TiO2) [27, 28], tantalum pentoxide (Ta2O5) [29, 30] and in very specific cases some proteins [2]. It must be considered that the deposition of these additional materials, increase the production costs of the proposed device and decrease the performance of the projected components, because the generated evanescent wave is hindered by this type of procedure, specially in integrated photonics.
In this work, the biofunctionalization process of the polymeric surface is developed directly on the surface of the polymeric material. In integrated photonics the biofunctionalization on an organic material that also works as transport medium for the light in our transducer, is an excellent alternative, due to the high affinity demonstrated in the proof of concept presented.
The demonstration of biofunctionalization presented in this manuscript and the procedures previously described provide a complete, suitable and low-cost platform for polymeric optical biosensor fabrication with simple surface treatment for the advancement of Lab-on-Chip (LOC) technologies.
2. Results and discussions
In order to use a conventional SPR system for characterization we start with a glass substrate covered with a 50 nm thick gold film cleaned in piranha solution (1:3 mixture of 30% H2O2 and concentrated H2SO4) for approximately 3 minutes, followed by immersion in acetone (5 minutes) and then in isopropyl alcohol (5 minutes). Afterward, the sample was fully washed with deionized water and dried with N2 flow. The selected photo-resist SU-8 2100 was diluted from 75% solid down to 3% solid with cyclopentanone and, then adjusted to weight concentrations of 0.1% H-nu 470 photoinitiator, 0.1% AN-910E cationic cure accelerator, and 2.5% OPPI photoacid generator (4-(octyloxy) phenylphenyliodium hexafluoroantimonate) [9]. The sample was spin coated with the modified phot-resist for 10 seconds at 500 rpm, followed by 30 seconds at 8000 rpm. The sample was then prebaked for 1 minute at 95 °C and exposed using a DLW system at 405 nm wavelength. After exposure, the photo-resist was subjected to the post-exposure bake for 1 minute at 95 °C. The polymerized surface was developed by submersion and stirring for 60 seconds. Finally, it was hard baked for 5 minutes on a hot plate at 150 °C. An illustration of the fabrication process can be seen in Fig. 1(a), while Fig. 1(b) shows the thickness curve variation as function of the spin coating speed (second spin step).
Fig. 1 Sample preparation process. (a) Main steps in the fabrication process of the thin SU-8 layer on gold. (b) Spin-coating characterization for spin speeds between 1000 rpm and 9000 rpm.
An innovative process was implemented using a DBD plasma for the hydrophilization of SU-8 surface, allowing the sample exposition to plasma treatment at room temperature in atmospheric pressure, ensuring the integrity of the exposed structures. This procedure is conducted using helium gas and high voltages in very short pulses. The DBD plasma process is considered a low cost procedure compared to others systems, because it does not need expensive vacuum assemblies and it does not limit the size of wafers. Besides, DBD plasma process does not leave toxic species over the SU-8 surface, which can cause difficulties under biofunctionalization [13] From the chemical structure of SU-8, after plasma exposure it is easier to break a C-C bond with binding energy of 284.8 eV on the surface than C-O or C-O-C bonds with binding energies of 286 eV. By breaking C-C bonds there is a greater probability of C-H bonds occurring; in fact this was observed via X-ray photoelectron spectroscopy (XPS) of the surface after the plasma treatment. We characterized the DBD plasma effect for different exposure periods by evaluating the surface contact angle reached and its subsequent deterioration. Maximal angle reduction was observed with 8 minutes of sample exposure, driving the angle from 80° down to 5° as can be seen in Fig. 2.
Fig. 2 Characterization of the effect of plasma exposure on the surface of the modified SU-8 as function of the time. Measured contact angle for different exposure intervals of DBD plasma treatment and recovery curve for the contact angle after 2 minutes exposure.
It is possible to appreciate decreases around 20° in less than 1 minute of exposure time, 8 times less than with oxygen or argon plasma, reaching contact angles below 10°, after only two minutes of exposure time. In addition, Fig. 2 shows the recovery curve for the contact angle as a function of aging time, where it is possible to observe slow recovery during the first 2 hours with contact angle below 10°, and faster recovery afterwards. Even after almost 80 hours of testing the initial contact angle was not recovered, ensuring good surface hydrophilicity during the whole experiment.
Surface roughness is also of paramount importance of optical applications, being fundamentally responsible for most of the scattering losses in transparent materials, such as SU-8. In Fig. 3, a study of RMS surface roughness is carried out for three different samples: before plasma treatment we observe a value of (0.20 ± 0.05) nm, after plasma exposure for 2 minutes the roughness value increased to (0.40 ± 0.012) nm, and after 10 minutes to (1.80 ± 0.3) nm. All measurements were performed by atomic force microscopy (AFM). This small roughness increase as function of plasma exposure time has not been studied before with radio frequency (RF) plasma treatments [15, 16], and, together with the contact angle results, represents an advantage for the proposed DBD plasma process for hydrophilization of organic surfaces. Nonetheless, after 10 minutes exposure, macroscopic irregularities appear on the surface of the samples. Because proper hydrophilization occurs in much shorter periods (as seen in Fig. 2), the slight increase in roughness does not compromise the performance of our surface during the detection process or in future integrated optical devices.
Fig. 3 SU-8 + photoinitiator surface characterization obtained by AFM. (a) Surface roughness of the SU-8 + photoinitiator layer without DBD plasma exposure. (b) Surface after exposure for 2 minutes. (c) Surface after exposure for 10 minutes.
We used SPR techniques to evaluate the efficiency of the nano-layer of SU-8+photoinitiator after treatment as an optical biosensor. A surface plasmon can be described as the collective oscillation of valence electrons in a solid by incident light. The resonance condition is established when the frequency of the light photons matches the natural frequency of surface electrons oscillating against the restoring force of positive nuclei [31–33].
This measurement system is composed by two lasers working at 670 nm and 785 nm, respectively. The lasers are directed to a prism described in Fig. 4(a), allowing the generation of the plasmon in the metal surface seated in the flat section of the prism. The prism, the gold plate and the polymeric layer are positioned on an articulating base, controlled by a high-precision step motor that allows controlling the angle of incidence of the light. In this way, when the frequency of the incident light coincides with the plasma resonance frequency, a decrease in the reflected / transmitted energy (resonant absorption) will be seen at any particular incident angle.
Fig. 4 Polymer-based surface plasmon resonance. (a) Simulated light behavior in SPR detection system. (b) Comparison between simulated and experimental resonance in SU-8-based SPR detection system.
In order to guarantee that the electronic surface plasmon to exist, the real part of the dielectric constant of the metal must be negative and its magnitude must be greater than that of the dielectric. This condition is met in the visible wavelength for metal/air and metal/water interfaces. Taking advantage of this condition, we have deposited a modified SU-8 layer of only 25 nm of thickness, due to the small thickness of the polymeric nano-layer the light is not completely confined, generating an extended evanescent wave in the the highest refractive index medium, in this case, the external environment, providing a transition environment for the evanescent wave from the metal (gold) layer with 50 nm of thickness, to the polymer surface, generating an evanescent wave with approximately 1 µm depth in the external medium reaching Limit of Detection LOD) of the order of 10−5 in our SPR systems, as can be seen in Fig. 4(b).
To verify this, we have performed numerical simulations of an SPR with incident angle at 51° composed by four different media: BK7, gold, SU-8+photoinitiator and air. In Fig. 4(b) a discontinuity can be observed which represents the transition of the light from the metal to the polymer layer, where the bioreceptors will actually be immobilized. In Fig. 4(c) the variation of the measured and simulated reflectance as function of the variation of the angle of incidence of light is presented. Analyzing the response of the device at 670 nm and 785 nm wavelengths, good agreement between experimental and simulated results is observed, with a small difference in the curve depth simply because material losses were not included in the simulations.
Immediately after plasma exposition, the terminal carboxyl groups on the SU-8 surface were activated by the addition of an aqueous solution containing N-(3-Dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC) (150 mmol/L) and N-Hydroxysuccinimide (NHS) (100 mmol/L) for approximately 15 minutes for the formation of the NHS-ester groups.
Two different biomolecules were immobilized for the construction of biosensors. The first one was the standard protein IgG to evaluate its association and dissociation on polymeric surface previously activated, as shown at Fig. 5(a). The second one was the Prostate Specific Antigen (PSA) in order to evaluate a specific system immobilization, which was used as a recognition element in Fig. 5(b), for the detection of anti-PSA in Fig. 5(c). These results give evidence to the sensorgramms (ΔΘSPR vs time) where three typical regions of the curves for each one of the biomolecules can be observed.
Fig. 5 Immobilization of the IgG and Prostate Specific Antigen (PSA) on SU-8 activated surface. (a) Sensorgramm dynamic in response to the association or dissociation of proteins. (b) Same as (a) but for PSA. (c) Response to the interaction antigen-antibody after anti-PSA deposition.
In Fig. 5(a) and Fig. 5(b), the highlighted regions are very similar. In region one, it is possible to observe a baseline obtained from the addition of HBS-EP buffer solution at pH 7.4 on the previously activated SU-8 film on gold surface (SU-8/Au). The addition of IgG in Fig. 5(a), or the immobilization of the PSA in Fig. 5(b), is represented by region two, where variations of the SPR angle (ΔΘSPR) for each concentration of IgG (100 and 250 µg/mL) or PSA (2 µg/mL) added on polymeric surface were observed. This transition occurred because the binding of the protein on the SU-8/Au increased the local refractive index in the proximities of the polymeric surface, which led to alteration in the plasmon resonance condition. It is easy to visualize from Fig. 5(a) and Fig. 5(b) that the equilibrium is quickly reached in approximately 15 minutes for both cases. In region three we see the response of our system to successive additions of HBS-EP buffer solution at pH 7.4 for the removal of weakly adsorbed biomolecules on the SU-8/Au. The dissociations for IgG show a reduction in ΔΘSPR of 1.06% and 1.01% for concentrations of 250 µg/mL and 100 µg/mL respectively, suggesting strong binding of the biomolecule to the proposed SU-8+photoinitiator film. For PSA the reduction was 22% for a concentration of 2 µg/mL.
After the bond of PSA, the unbound reactive NHS-ester groups were deactivated by a brief flow of 1 mol/L of an ethanolamine aqueous solution for approximately 5 minutes to prevent non-specific binding. The excess of unbounded ethanolamine molecules was removed through several additions of water. After this, the SPR slide sensor was dried with N2 flow. Finally, the evaluation of the antigen-antibody interaction was carried out by the addition of anti-PSA antibody (1 µg/mL). Fig. 5(c) gives evidence to the association (region two) and dissociation (region three) phases. It is possible to observe that the addition of anti-PSA in very low concentration (1 µg/mL) was accompanied by a significant response, which suggests the high potential of the proposed material (SU-8 film with photoinitiator H-nu 470) for biosensing applications.
3. Conclusion
The fabrication of a few nanometer thick film of SU-8 + H-nu 470 photoinitiator by direct laser writer, hydrophilization treatment with DBD plasma and a proof of concept of biodetection procedure through an SPR-based polymeric system for standard and specific biomolecules were presented. As a consequence of the biofunctionalization carried out directly on the polymeric surface, the efficient implementation of an organic nano-layer in a SPR measurement system was demonstrated. The proposed exposure of the polymer surface to DBD plasma provides surface with approximately 5° of contact angle and roughness around (0.40 ± 0.012) nm. As proof of concept, it was possible to observe a dissociation of only 1% after removal of weakly adsorbed standard biomolecules on the SU-8/Au. In addition, the biodetection of a specific antigen with great performance at low concentrations was also demonstrated, showing that the surface of SU-8 + H-nu 470 provides a reliable and stable platform with great potential to act as a biosensor device.
Funding
Brazilian MCTI FOTONICOM (CNPq project no. 574017/2008-9 and FAPESP project nos. 2008/57857-2 and 2009/54045-0); INCTBio (465389/2014-7).
Acknowledgments
The authors would like to thank to Dr. Fellype do Nascimento, Prof. Dr. Stanislav Moshkalev and Prof. Dr. Munemasa Machida from University of Campinas for the fruitful conversations and support in the development of some key experiments for our final conclusions; to thank the Information Technology Center Renato Archer (CTI), the Center for Semiconductor Components(CCS) and the Plasma Laboratory of the University of Campinas for making available the required resources for the development of this work; and finally to thank the Brazilian MCTI FOTONICOM (CNPq project no. 574017/2008-9 and FAPESP project nos. 2008/57857-2 and 2009/54045-0)and INCTBio project no. 465389/2014-7 for the financial support.
Disclosures
The authors declare that there are no conflicts of interest related to this article.
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