Introduction

The extraordinary optical transmission (EOT) through subwavelength nanoholes built on thin films of gold and silver was first reported by Ebbesen and coauthors in 1998 [1]. They observed a greater light transmission than what was predicted by the classical theory of Bethe on light diffraction through holes with dimensions of the order of the light wavelength [2]. The EOT is related to the surface plasmon resonance (SPR) excitation, which involves trapped light waves on the metal surface due to oscillation in resonance with electrons of the metal conduction band [1, 3]. The SPR phenomenon is highly sensitive to variations in the refractive index on the metal surface that supports surface plasmons (SPs). Therefore, plasmonic devices based on nanohole arrays constructed on gold films have been used to investigate binding events that occur on the metal surface in affinity biosensors while performing optical measurement [4, 5]. A SPR affinity sensor is characterized by a recognition element (immobilized on metal) that can recognize and interact with a specific analyte of a sample that is in contact with the metallic surface of transducer; this, in turn, converts the event binding in a measurable optical signal [3]. These sensors usually evaluate the band wavelength shift of the transmission spectrum in response to the molecular recognition of the analyte [4–6]. However, an investigation of changes in transmitted light intensity can also be employed [7, 8].

Traditional SPR biosensors based on the Kretschmann configuration require a prism attached to a thin metal film for excitation of SPs [9]. This type of sensor has been employed in various bioassays that range from the detection of analytes related to food quality and increased safety [10–14] to environmental monitoring [15–21]. This type of sensor has also been used for medical diagnosis regarding the detection of drugs [22–24], hormones [25–27], and markers for cancer [28–35]. Despite the wide application and high sensitivity of the conventional SPR configuration, these devices have difficulty of setup alignment, which is a great experimental inconvenience [36]. Furthermore, these devices are not compact, which is an inconvenient feature that limits their application to field analysis and the degree of multiplexing. On the other hand, nanohole array-based plasmonic sensors overcome these experimental drawbacks. There is the possibility of operating in collinear optical arrangements with minimal sensory areas, which facilitate integration with microfluidic systems, multiplexing, and miniaturization [37, 38]. Nanohole array-based plasmonic biosensors built on gold thin films have been applied to biomolecular detection at various levels [4, 38–40].

The determination of tumor biomarkers is of great interest in medical diagnostics. These analytes are used to indicate the presence of cancer in patients, predict the behavior of a disease in response to treatment, predict the chance of patient recovery, and even help detect cancer in apparently healthy or high-risk people before the symptoms appear [41]. For instance, the detection of the paired box gene 8 (PAX8) protein, an important biomarker of ovarian cancer, has been reported by using nanohole array-based SPR biosensors on thin metallic films [39, 40].

Considering the advantages and applications of plasmonic biosensors based on nanostructured metallic films for traditional SPR devices, in this paper, we propose the detection of human epidermal receptor protein-2 (HER2), a cancer antigen, by using a biosensor based on a gold nanohole array integrated into a microfluidic system operating in spectral mode. The HER2 antigen is an important cancer marker that is in 20–25 % of individuals with breast cancer [42–45], and its determination is interesting since it is a marker that is used for the diagnosis and prognosis of the disease [45–47]. It is noteworthy that to our knowledge, the determination of this antigen by using SPR biosensors based on nanostructured gold film was not previously reported.

Experimental

Obtaining the Substrate Containing Nanohole Arrays

Square glass slides (2.5 × 2.5 cm and 1.0-mm thick) covered with 100-nm-thick gold films were used. A thin chromium layer (5-nm thick) was used as an adhesive layer between the glass and gold. These substrates were employed to obtain the nanohole arrays by a focused ion beam (FIB, FEI 235) with a dual-beam device by using a gallium ion source. The ion beam was set to 30 keV with a milling rate of 1.6 nm ms−1, and the beam current was 300 nA. The diameter of each nanohole was 200 nm, and array periodicity was 400 nm. The substrates were characterized by scanning electron microscopy (SEM, Shimadzu Superscan SSX-550).

Microfluidic Chip

The microfluidic system was obtained by using well-established soft lithography techniques [48]. Square glass substrates (2.5 × 2.5 cm and 1.0-mm thick) were cleaned with acetone and piranha solution H2SO4 (FMaia, 95 %)/H2O2 (Anidrol, 35 %) in a 3:1 (v/v) ratio. Photomasks were created and printed on polyterephthalate sheets by using high-resolution printing. The photoresist (SU-8® 50, Microchem) was spin coated (Spincoating Systems, G3P-8) on a glass slide at 500 rpm for 5 s, dwelling for 8 s, ramping at 2000 rpm for 5 s, and spinning for 25 s. This resulted in a uniform thickness of ca. 50 μm. The coated glass slides were prebaked at 65 °C for 6 min and then at 95 °C for 20 min to harden the photoresist. The photomask was positioned on top of the photoresist film, and it was exposed under collimated UV light (Tamarack Scientific, 2110CP) for 17.5 s. The sample was then baked, first at 65 °C for 1 min and then at 95 °C for 5 min to promote cross-linking of the photoresist. The exposed glass slide was immersed in SU-8 developer until the unexposed photoresist was completely removed, which created the master templates. The microchannel structures were created by curing a solution of polydimethylsiloxane (PDMS, Sylgard 184®) and curing agent (Sylgard 184®) at a ratio of 10:1 w/w in order to create PDMS on the master templates. The PDMS was then degassed for 30 min in a vacuum and then cured at 95 °C for 2 h. The thickness of the PDMS was ca. 4.0 mm and microchannels with a width of 70 and 350 μm in spacing between channels.

Solution Flow and Optical Measurements

To obtain the transmission spectra, an experimental arrangement was used in a collinear mode, as shown in Fig. 1. This arrangement used a halogen lamp (300 W) with a broad emission band. The light was perpendicularly (normal) focused and collimated to the metallic nanohole arrays surface by using a ×50 objective lens (Meiji Techno, WD 7.5 mm) coupled to a trinocular microscope. The transmitted light was collected by an optical fiber (Ocean Optics, P600-2-UV-vis) positioned under the array. This optical cable was coupled to a UV-vis spectrophotometer (Ocean Optics, USB2000), and thus, transmission spectra were recorded and stored into a computer. The nanostructured metallic substrate was integrated with a microflow cell by using a PDMS piece that contained microchannels aligned to the array, as shown in the inset of Fig. 1. The flow of solutions was carried out with the aid of a syringe pump (Harvard Apparatus 11 Plus).

Fig. 1
figure 1

Experimental optical arrangement integrated with microfluidic system in collinear transmission mode to obtain spectra of transmission across substrates that contain nanohole arrays on gold film. Inset: photographic image that shows the alignment of the PDMS microchannel with the array

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Sensor Sensitivity Characterization

The optical response of the sensor to changes in refractive index was performed by acquiring transmission spectra of the nanohole arrays in solutions with different refractive indices. For this purpose, aqueous solutions of d-(+)-glucose (Sigma, 99.5 %) at concentrations of 2.0, 6.0, 10.0, 14.0, 18.0, 22.0, 26.0, and 30.0 g 100 mL−1 were used. Each solution had its refractive index determined (with 0.0001 precision) by using a portable digital refractometer (Atago, 3850 PAL-RI).

Biosensing Analysis of HER2

An appropriate strategy was used to immobilize the molecular biorecognition element (antibody against HER2) on the metallic nanostructured surface. For gold surfaces, the immobilization is normally performed via streptavidin immobilized previously on biotin [49–51]; in turn, the biotin is attached to the surface via covalent interaction by using an alkanethiol organized monolayer on the gold surface [52].

The surface with the biorecognizing element that would be received was prepared initially by immersing the nanohole arrays in an ultrasonic ethanolic bath for 5 min. Soon after, it was extensively rinsed with water and dried with N2, and then, it was integrated with the microfluidic system. Subsequently, an aqueous solution of cysteamine (Aldrich 95 %) 0.06 mol L−1 was flowed on the substrate surface for 72 h. Afterwards, a 5-mg mL−1 aqueous solution of biotin-sulfo-NHS (Sigma-Aldrich, 90 %) was injected for 12 h. Then, 0.5 mg mL−1 of streptavidin in phosphate buffer solution (PBS, pH = 7.4) was passed for 4 h.

The previously prepared substrate was then modified by immobilization of biotinylated antibody against HER2 (AB1, Invitrogen, 95 %) via specific interaction with the previously immobilized streptavidin. This AB1 immobilization step was achieved by flowing a 50-μL mL−1 AB1 solution for 12 h. Soon after, a bovine serum albumin (BSA) solution (100 ng mL−1) was flowed on the substrate for 2 h. This last step was necessary to block nonspecific binding sites. Finally, the HER2 antigen (Invitrogen, 70 %) was flowed on the substrate surface (already containing the antibody) by using a 30-ng mL−1 antigen solution for 1 h. To use the sandwich detection mode, a 5-μg mL−1 solution of a secondary antibody against HER2 (AB2, Invitrogen 75 mg L−1) was flowed on the system for 2 h. The antibodies, the BSA, and the antigens were diluted in PBS.

After each modification step, the system was purged at a flow rate of 2 mL h−1 with PBS to remove loosely adsorbed molecules and the transmission spectra acquired. The flow rate of the successive modifications was 0.05 mL h−1.

Results and Discussion

The square gold nanohole arrays constructed by FIB had a periodicity of 400 nm and a hole diameter of 200 nm with an active area of ca. 14 × 14 μm, as shown in Fig. 2.

Fig. 2
figure 2

SEM images at different magnifications of a square nanohole array on gold film with 400-nm periodicity, hole diameter of 200 nm, and depth of 105 nm (thickness)

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Figure 3 shows the transmission spectra of the nanohole arrays immersed in glucose solutions of different refractive indices. Two bands were observed, and both undergo redshift as the refractive index increases, which is an expected behavior for plasmonic substrates [4, 53]. According to Eq. 1, the two bands are both related to the (1,0) modes for metal/liquid ((1,0)liquid mode) and metal/glass ((1,0)glass mode) interfaces, respectively. The \( {\lambda}_{\max } \) is the maximum wavelength of the plasmon, \( {a}_0 \) is the periodicity of the array, \( i \) and \( j \) are integers relative to the plasmonic modes, and \( {\varepsilon}_{\mathrm{d}} \) and \( {\varepsilon}_{\mathrm{m}} \) are the dielectric constants of the dielectric materials (\( {\varepsilon}_{\mathrm{glass}}=2.13 \) and \( {\varepsilon}_{\mathrm{liquid}}=1.78 \), in the oscillation frequency for visible light) and metal (\( {\varepsilon}_{\mathrm{Au}} \) depends on the wavelength and can be obtained elsewhere [54]), respectively.

Fig. 3
figure 3

Transmission spectra of nanohole arrays as a function of the refractive index (1.3344 to 1.3834). Array has a periodicity of 400 nm, and the hole diameter is 200 nm. The bands represent the (1,0)liquid and (1,0)glass plasmonic modes related to gold/liquid and gold/glass interfaces, respectively

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$$ {\lambda}_{\max }=\frac{\alpha_0}{\sqrt{i^2+{j}^2}}\sqrt{\frac{\varepsilon_{\mathrm{d}}{\varepsilon}_{\mathrm{m}}}{\varepsilon_{\mathrm{d}}+{\varepsilon}_{\mathrm{m}}}} $$
(1)

The (1,0)glass band possesses a higher sensitivity since the transmission maximum wavelength (λ max) shifts, at the same solutions, are higher for this plasmonic mode. As can be seen in Fig. 3, a redshift of ca. 9 nm was obtained for the (1,0)liquid mode, while for the (1,0)glass mode, the redshift was ca. 15 nm when the refractive index changed from 1.3344 to 1.3834.

Since the plasmonic (1,0)glass mode show a higher sensitivity, the λ max for that band was plotted as a function of the refractive index. The resultant curve slope (Fig. 4) is defined as the sensitivity of the plasmonic substrate, and its characterization is important in determining the ability of the substrate to have its plasmonic transmission band disturbed due to changes in refractive indexes on the metal surface [55]. The sensitivity obtained was ca. 305 nm RIU-1 (RIU = refractive index units) and is comparable to results reported in the literature for plasmonic devices of the same nature [5, 6]. It was also observed that there was a good linear fit since the linear fit coefficient (R 2) value was near unity (R 2 = 0.9976). It must be considered that a proportional response to the variation of refractive index is a sought parameter for SPR sensors. The inset of Fig. 4 shows the (1,0)glass region (magnified and normalized) obtained for each refractive index, which shows the redshift as the refractive index increased.

Fig. 4
figure 4

Curve of λ max shift as a function of refractive index for bands relating to (1,0)glass mode. The inset shows the (1,0)glass bands (expanded and normalized) for each refractive index tested. In the “a” to “h” range, the refractive index changes from 1.3344 to 1.3834

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Considering that the sensor proved to be able to detect variations in refractive indexes near the metallic surface, we used the sensor to test the detection of HER2, an antigen associated with breast cancer. For this, the nanohole gold surface was initially prepared by immobilization of a cysteamine monolayer and thereafter with biotin and streptavidin.

For the detection of the HER2 antigen, the surface received the AB1 molecules (biotinylated antibody) through its interaction with previously immobilized streptavidin molecules. Then, the HER2 antigen was immobilized, and finally, the surface received the AB2 antigen. The use of a secondary antibody, which characterizes the sandwich detection mode, improved the sensitivity of the sensor [3]; the antigens are molecules of relatively low molecular mass (small volume), which does not produce significant changes in refractive index. It is often difficult to detect the antigen by spectral monitoring. However, antigen conjugated with bulky secondary antibody causes a relatively larger change in the refractive index of the metal, which is then more easily detected through a spectral band shift. The surface modification scheme (starting from surface prepared with the cysteamine-biotin-streptavidin) is shown schematically in Fig. 5.

Fig. 5
figure 5

Scheme of HER2 antigen detection via immobilized biotinylated antibody (AB1), starting from a metal surface that was previously prepared through a cysteamine streptavidin-biotin process. For the sandwich detection mode, a secondary antibody (AB2) was used

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Figure 6 shows the transmission band shift for the (1,0)glass mode during the steps of the HER2 antigen detection (starting from AB1 immobilization).

Fig. 6
figure 6

The maximum wavelength shift (λ max) for the (1,0)glass mode as a function of the metallic surface modification steps. The AB1 antibody immobilization was taken as zero, and the HER2 antigen concentration detected was 30 ng mL−1. Inset: normalized and expanded transmission (1,0)glass band spectra for each modification step

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The spectral shift observed for the (1,0)glass band (Fig. 6) is expected. As the molecules are immobilized on the surface, the refractive index on the metal/dielectric interface increases, leading to a redshift. Compared to the AB1 immobilization step, a shift of ca. 1.2 nm was observed when the HER2 antigen (concentration 30 ng mL−1) was immobilized. An extra redshift of ca. 1.3 nm was observed when AB2 was linked to the AB1/HER2 layer; this resulted in a 2.5-nm total band displacement. This shift is comparable with those obtained in other studies that report the use of sensors based on nanohole arrays applied in biomolecular detection [56, 57]. Furthermore, considering that the transmission maximum shift obtained for the immobilization of AB2 was of 2.50 ± 0.03 nm (with respect to surface covered with AB1) and the fact that our system is capable of measurements with an accuracy of 0.1 nm, the signal-to-noise ratio is high enough so that the sensor can be efficiently used to detect HER2 antigen (in the tested concentration) and the observed shift is easily measured by our system.

We observed that the shifts regarding the immobilization of HER2 and AB2 are similar (1.2 and 1.3 nm, respectively). However, AB2 molecules are larger than the HER2 ones, which should imply a higher plasmonic-disturbing effect, but this was not observed. This possibly occurs because the SPs are confined on the metal surface, and the amplitude of the plasmon electromagnetic field (and hence sensitivity) decays exponentially as it moves away from the metal/dielectric interface [58]. Thus, the AB2 molecules cause a smaller plasmonic perturbation than expected because they are relatively farther away from the surface (regarding the antigen molecules). Nevertheless, the sandwich detection mode clearly provided a greater sensitivity to the presence of biomolecules on the metal surface.

Conclusion

The nanohole array-based microfluidic sensor presented in this work was sensitive to the presence of the breast cancer biomarker at a concentration of 30 ng mL−1. It is noteworthy that HER2-positive individuals have concentrations between 15 and 75 ng mL−1 in blood plasma [59]. Thus, the sensor shows potential to be applied in the diagnosis or prognosis of breast cancer. It is interesting to note that the SPR substrate showed good sensitivity (ca. 302 nm RIU−1) and a great linear behavior when compared to other nanohole array-based plasmonic biosensors for spectral investigation.