A Novel Thermal Bubble Valves integrated Nanofluidic

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A Novel Thermal Bubble Valves integrated Nanofluidic Preconcentrator for high sensitive biomarker detection Chih-Zong Deng, Yu-Jui Fan, Pei-Shan Chung, and Horn-Jiunn Sheen ACS Sens., Just Accepted Manuscript • DOI: 10.1021/acssensors.8b00323 • Publication Date (Web): 11 Jun 2018 Downloaded from http://pubs.acs.org on June 11, 2018

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A Novel Thermal Bubble Valves integrated Nanofluidic Preconcentrator for high sensitive biomarker detection Chih-Zong Deng†, ‡, Yu-Jui Fan‡,§,*, Pei-Shan Chung∥ and Horn-Jiunn Sheen†,* †

.Institute of Applied Mechanics, National Taiwan University, No. 1, Sec. 4, Roosevelt Road, Taipei 10617, TAIWAN, R.O.C. ‡

.School of Biomedical Engineering, Taipei Meidical University, 250 Wuxing St. Taipei 11031, TAIWAN, R.O.C.

§

.International Ph.D. Program for Biomedical Engineering, Taipei Medical University 250 Wuxing St. Taipei 11031, TAIWAN, R.O.C. ∥

.Department of Bioengineering, University of California at Los Angeles, 420 Westwood Plaza, Los Angeles, CA 90095, USA *E-mail: [email protected]; [email protected].

KEYWORDS: nanofluidic preconcentrator, acoustic wave, oscillating thermal microbubble, immunobeads, Brownian motion, C-reactive protein

ABSTRACT: In this study, we developed a new immunosensor that can achieve an ultra-low detection limit and high sensitivity. This new device has an electrokinetic trapping (EKT)-based nanofluidic preconcentrator, which was integrated with oscillating bubble valves, to trap concentrated antigen and immunobeads. During the immunoassay process, oscillating bubbles rapidly grew and acted as control valves and to block the microchannel. Thereafter, the trapped preconcentrated antigen plug and antibody-coated nanobeads were preserved in the region between these two valves. Finally, the antigen concentration was quantitatively analyzed by a real-time measurement of Brownian diffusion of the immunobeads. In this work, the test sample used was C-reactive protein (CRP) which is a risk indicator of coronary heart disease and atherosclerosis.

In the past two decades, microfluidic and lab-on-a-chip technologies have been extensively used in many areas such as analytical chemistry, diagnosis of diseases, and the pharmaceutical industry.1-2 One of the major challenges in biomolecular detection is the sensing of analytes with low abundances. A preconcentrator is an important component to raise the concentration of a sample beyond the detection limit of biosensors. A number of studies were carried out, including field-amplified sample stacking,3 isotachophoresis,4 porous silica membranes,5 and ion concentration polarization (ICP).6-8 An ICP-based preconcentrator was developed which employs a nanofluidic filter as an ion-selective membrane to cause a sample to accumulate at the boundary of the nano/microfluidic channel by electrokinetic trapping (EKT). Han et al. used Nafion® membranes as a nanofluidic filter to create a high potential preconcentrator, which was simply fabricated, was suitable for any charged analyte, and had a high preconcentration factor.9-10 A preconcentration factor of up to a million-fold was achieved.11 However, the EKT-based micro/nanofluidic preconcentrator has a major challenge which is due to unstable electrokinetic flow in the microfluidic channel. The unstable flow disturbs the antibody-antigen interaction when the

immunoassay is carried out. To overcome this issue, some research integrated pneumatic valves into the nanofluidic preconcentrator9, 12-13 to isolate the concentrated antigen and antibody-modified reacting area for immunosensing. However, in order to fully block the channel, the pneumatic valve needs to be operated using high-pressure air. This approach caused system miniaturization to become difficult. The use of surface acoustic waves to oscillate microbubbles has attracted much interest for its versatile applications including as a pump, mixer, filter, and transporter.14-20 Microbubbles can be generated by two mechanisms: (1) thermal-accumulated liquid-gas phase change, and (2) geometrically trapped air bubbles while flowing through a microfluidic channel. In the former, when the liquid phase changes to vapor phase, the explosively expanding volume can be used as a pushing force for liquid injection, particle shift for a cell sorter, a droplet generator, and a cell membrane cutter.21 Normally, thermal bubbles are not stable due to thermal dissipation. In the second case, geometrically trapped air bubbles can be simply generated by changing the microfluidic channel geometry, i.e., into a horseshoe-shaped structure. The gas-liquid interface can be driven by an acoustic wave which dis-

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turbs the fluid.18 Geometrically trapped bubbles are stable when the channel is at room temperature. Huang et al.22 demonstrated a chemical waveform switch using different sizes of microbubbles, which can be selectively driven by acoustic waves of different frequencies on the same microchip. C-reactive protein (CRP), a sensitive biomarker of inflammation, is an important risk factor for atherosclerosis and heart disease. CRP concentration level will reach up to three to four orders rise when people got inflamed.23 Currently, ELISA is generally used for the clinical test, but may take long time for multiple-step reaction.24 Some other major label-free diagnostic approaches, e.g. surface plasmon resonance (SPR)25, quartz crystal microbalance (QCM)26, and cantilever beam27, may also be the major probable methods for rapid detection. In this research, we developed a new application of acoustically excited microbubbles as a control valve, which was further integrated with a nanofluidic preconcentrator to trap a collecting plug. Since we would like to generate microbubbles in a microfluidic channel, the bubble size and generating location were our channel design concerns. We also demonstrated that by selecting the acoustic wave frequency, which is due to the bubble size, we could turn on or off the bubble valves. The region between the two bubble valves was isolated for antibodyantigen interactions. The bubble valve-integrated preconcentrator was also previously developed to enhance the sensitivity of the technology to immunosense the nanobeads’ Brownian movement.28 The experimental results showed that the limit of detection could be improved by 4 times order-fold. In comparison to current state of the art methods, this approach can be applied to measure samples’ CRP concentration level with and without preconcentration and further to double check the measurement data.

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the microchamber, was 10 times smaller than that of the microchamber. The ratio of 10:1 was sufficient to stably maintain the concentrated plug inside the microchamber. As shown in Figure 1, the microchamber with a width of 240 μm was designed to match the field of view of the camera. The channel width is 26 μm. Four triangular sidewall cavities, each has a height of 210 μm and a base of 40 μm, are used for stably trapping bubbles. The smaller the ratio of the cavity base to the channel width is, the more stable is for trapping the bubble in the cavity. However, a smaller bubble has a higher resonant frequency that makes the piezo-transducer possible to be overheated and damaged. This is a trade-off so that we designed the ratio of the cavity base to the channel width was 1.5 to provide stable bubble trapping and to prevent the piezotransducer from damaged. Two pairs of sidewall cavities were used to trap bubbles when the microchannel was filled with phosphatebuffered saline (PBS). An acoustic piezoelectric transducer was attached to the device to generate acoustic waves.

Experiments Design of the microchip. Figure 1 shows a schematic diagram of the microchip. The preconcentrator was made of a main channel and a buffer channel. Both channels were bridged by a thin Nafion® film, which was used as an ion-selective channel. The preconcentration mechanism, the channel design, and operational procedures were described in our previous study.29 Based on this design, a concentration plug can be formed on either side of the Nafion® film in the microfluidic channel depending on the applied voltage difference. The microchamber, right beside the Nafion® filter, was used to collect and to isolate the concentration plug. The width of the microchannel, which connects to

Figure 1. Schematic diagram of the preconcentrator with a bubble valve.

Microchip fabrication. The single-layer microchannels, consisting of a main channel and a buffer channel, were fabricated by a general Polydimethylsiloxane (PDMS) fabrication process. First, the master PDMS microchannel on a silicon wafer

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was fabricated by standard photolithography using SU-8 2025 photoresist. The thickness of the photoresist on the silicon wafer was 25 µm. Then the PDMS base and curing agent were mixed in a 10:1 weight ratio. The mixture was poured onto the master after degassing for 1 h. The PDMS microchannel was peeled off from the master after curing on a 65 °C hotplate for 4 h. The strip of Nafion® (70160, Sigma-Aldrich) was deposited on a glass substrate using a microflow surface-patterning method.10 Then the aligned PDMS microchannel and the Nafion® patterned glass substrate were bonded together after treatment with oxygen plasma. Finally, a piezoelectric transducer was attached to the glass substrate near the main channel with epoxy. Experimental procedures. Figure 2 illustrates the step-by-step procedure of capturing the concentrated plug in the microchamber. First, the sample solution, which was mixed with CRP- and anti-CRP-coated nanobeads in a PBS solution, was injected into the main channel, and the buffer channel was filled with the PBS solution. Then, initial voltages were applied at both ends of the main channel to induce a depletion region. Both grounds were given at the buffer channel. When the depletion region formed near the vicinity of the Nafion® membrane, the voltage at one side was increased to accumulate the CRP. The electrical potential difference of the main channel was increased so that the concentrated plug could be transported to the reaction chamber. Lastly, the piezoelectric transducer was turned on to generate an acoustic wave. A pair of bubbles was triggered to oscillate by the acoustic wave, and then the size was expanded until it could block the microchannel and capture the concentrated plug.

Figure 2. Schematic diagram of the process to capture the concentrated plug.

Results and Discussion Acoustically excited bubble valve.

shown in Fig. 3(a). When the acoustic wave was applied, the trapped microbubbles began to expand until the bubbles fully blocked the main channel as shown in Fig. 3(b). The total thermal energy, Etotal, in the air bubbles was equal to the difference in the acoustically oscillating thermal energy on the air-liquid interface, Eacoustics, and thermal energy dissipation by fluid flow, Edissipation. The schematic diagram of the air bubbles is shown in Fig. 3(c), and the equation is given as follows:

Etotal = Eacoustics − Edissipation

(1)

The parameters that affect the increased thermal energy are dependent on several factors, including the distance between the bubbles and the acoustic source, and the frequency and amplitude of the acoustic wave. When the microbubbles were closer to the acoustic source, they obtained more energy. Similarly, when the acoustic wave had a larger amplitude or was at the resonant frequency of the bubbles, the boundary of the bubble oscillated and quickly grew. Eventually, the bubbles blocked the channel.

Figure 3. Microbubble valve. (a) A microbubble is trapped in the side-wall cavity, and (b) the microbubble is blocking the microchannel by absorbing the acoustic wave energy. (c) Schematic diagram of the change in the thermal energy of the air bubble.

The microbubbles were trapped in the side-wall cavities when the main microchannel was filled with liquid as

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The liquid-air interface of the bubbles oscillated due to the acoustic wave, and reached the maximum oscillating amplitude when at the resonant frequency. The resonant frequency, f, can be approximately estimated by the small-amplitude behavior of the Rayleigh-Plesset equation:30

f2=

  2σ 3κ  p +  4 ρπ α   α 1

2

2

 2σ  −   α 

(2)

where ρ is the density of the liquid, α is the radius of the bubble, κ is the polytropic exponent for a bubble containing air, p is the fluidic pressure, and σ is the surface tension of the solution.

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would not grow, which means that the thermal energy absorption and relaxation by the bubble were dynamically equivalent. The breakdown voltage of this piezoelectric transducer was observed to be 22 V. Based on this result, a maximum operating voltage of 20 V was used in the study. The relationship of the flow rate in the microfluidic channel versus bubble expansion rate was studied, and results are shown in Fig. 4(c). From the experiment, we found that at a given voltage of 20 V and a resonant frequency of 93.5 kHz, as the flow rate reached 6.4 μL/min, the bubble expansion rate was reduced to zero because of dissipation of the thermal energy by fluid flow.

Equation (2) uses the assumption that the bubble is spherical; however, in the study, the microbubbles trapped in the side-wall cavity were approximately a circular sector shape. Therefore, there was a certain deviation between the actual resonant frequency and the calculated results from Eq. (2). Once a certain fluid is chosen, the side-wall cavity design will determine the radius of the bubble; finally, the resonant frequency can then be estimated. Based on the cavity design as shown in Fig. 3(a), the resonant frequency was estimated to be 83 kHz. As mentioned earlier, due to differences in bubble shapes, the resonant frequency of 93.5 kHz was experimentally obtained. To investigate the characteristics of the bubble valve, we carried out an experiment to analyze the expansion rate of microbubbles in different conditions. The results are shown in Fig. 4. Figure 4(a) shows that the bubble dramatically expanded when excited at the resonant frequency, a Vpp of 20 V, and a flow rate of 0.8 µL/min. When the operating frequencies were increased or decreased by 10 kHz from the resonant frequency, the expansion rate was around onethird of that when operated at the resonant frequency. The microbubbles did not grow to block the microchannel when the frequency of the acoustic wave was exerted at ≤85.5 kHz, because the heat dissipation rate through the fluidic flow was higher than the thermal energy accumulating due to the acoustic wave. The oscillating amplitude of the air-liquid interface of the bubble is related to the voltage exerted by the piezoelectric transducer. In our experiment, the microbubble’s expansion rate was linearly propositional to the voltage exerted by the piezoelectric transducer at the resonant frequency and a flow rate of 0.8 µL/min as depicted in Fig. 4(b). When the exerted voltage was 14 V, the microbubble

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ACS Sensors spectively. The two side-wall cavities had the similar geometry but with different dimensions. The expanding microbubbles in the two side-wall cavities can also individually function as control valves. From Eq. (2), the resonant frequency of the microbubble depends on its size at a certain fluid condition. To clarify this phenomenon, we designed different sizes of side-wall cavities to trap microbubbles. The microbubbles were excited at their resonant frequency after becoming trapped in the cavity. Figure 5 shows experimental results of the bubble valves at various frequencies. When the piezoelectric transducer was set to manipulate the larger bubble at a resonant frequency of 24 kHz, it grew until it fully blocked the 60 μm channel in 315 s. In the meantime, the smaller bubble did not grow, because dissipation of the thermal was equal to or greater than its accumulation. Similarly, when we applied the resonant frequency of 40.2 kHz to the smaller bubble, it grew until it had fully blocked the 30 μm channel in 20 s. However, the larger bubble did not expand. The blockage time was shorter as the smaller channel width diminished.

Figure. 4 Bubble expansion rates vs. various experimental parameters of (a) the frequency, (b) voltage, and (c) flow rate.

Figure. 5 Frequency-dependent bubble valves. (a) and (b) at working frequencies of 24 and (c) and (d) 40.2 kHz.

Frequency-dependent bubble valves.

Temperature of the biochip.

Another related test was conducted with two cavities, one larger and one smaller, to study the detailed functioning of the control valves. The resonant frequencies of the larger and smaller bubbles were 24 and 40.2 kHz, re-

The temperature of microfluidic chip is an important issue, especially since it is used to handle biomolecules, cells, tissues, etc. The heat accumulating due to vibrations of the piezoelectric transducer or oscillations of the bubble air-liquid interface might lead to an increase in the

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chip temperature. To clarify this thermal issue, a thermocouple was used to directly monitor the temperature of the microchamber, which is the biomarker sensing area. Figure 6 shows results of temperature measurements on the microchamber when the piezoelectric transducer was operated for 40 s and 10 min, respectively. The operating voltage was 20 V, and the operating frequency was 93.5 kHz, which were also used for subsequent immunoassay experiments. As shown in Fig. 6(a), the temperature of the microchamber only increased by 1.5 °C and then decreased to its initial state within 120 s. Furthermore, the temperature of the microchamber increased to 30 °C when the piezoelectric transducer was operated for 10 min. When the piezoelectric transducer was turned off, the temperature dramatically decreased to the initial value within 3 min as shown in Fig. 6(b). For practical applications, the operating time of the piezoelectric transducer was only 40 s. However, the maximum temperature in the microchamber was 30 °C when the piezoelectric transducer was operated for 10 min. Based on these results, the temperature of the microchamber remained lower than 26.4 °C which is suitable for biomarker detection. The temperature of 30 ℃ will not affect the protein properties to our knowledge. Further, the acoustic piezo-transducer was turned off after the bubble valves fully blocked the channel, so that there was no mechanical force exerted when biosensor measurement was carried out. In this study, we also observed that there is no bubble generated in the flow channel. When the trapped bubbles fully blocked the channel, there was no fluid flow in the microchannel. Therefore, there is no heat dissipation by the fluid flow. After the acoustic piezo-transducer was turned off, the bubble valves fully blocked the channel and the bubbles would not collapse.

Figure. 6 Temperature of the microchamber. (a) The piezoelectric transducer was operated for 40 s and (b) 10 min.

Trapping of the preconcentration plug. Figure 7 (a) shows a microscopic view of the preconcentrator, bubble-valves, and microchamber. Here, the preconcentrator includes the main channel, the Nafion membrance, and the buffer channel. The preconcentrating procedures are exhibited step by step in Fig. 7(b)-(f). In order to visualize the flow field, a fluorescent molecular dye was added to the PBS solution. First, the fluorescent solution and the buffer solution were separately injected into the main channel and buffer channel. Because the fluorescence intensity was too low, the fluorescent solution could not be observed. When 25 V was applied to the preconcentrator, an ion-depleted region was formed near the Nafion membrane, and then, a preconcentrating plug began to form in the main channel when the applied voltage on the right side of the channel increased as shown in Fig. 7(b). In order to manipulate the preconcentrating plug, we adjusted the applied voltages at both ends of the main channel. The preconcentrating plug was manipulated to move into the microchamber as shown in Fig. 7(d) and (e). In the meantime, we turned on the pie-

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zoelectric transducer to generate the acoustic wave. The trapped bubble began to oscillate and expand thus blocking the main channel. Figure 7(e) shows that the preconcentrating plug was trapped inside the microchamber. After a while, the preconcentrating plug began to diffuse as shown in Fig. 7(f).

our previous study23, was used to modify the anti-CRP onto the nanobeads. Use of antibody-coated nanobeads for an immunoassay provides a dramatically enlarged reaction area for detecting targets compared to immunoassays carried out at the substrate surface. Because the nanobeads randomly moved in three dimensions, this phenomenon led to higher probabilities for conjugation of the antibody with the antigen. The random motion of particles suspended in fluids is called Brownian motion, which depends on the particle size, temperature, and viscosity of the solution. In 1905, Albert Einstein combined Stokes’ law to statistically analyze Brownian motion. The self-diffusion coefficient, D, for a dilute suspension given by Stokes-Einstein is:

D =α

kT 3πµ d p

(3)

where α is the correction factor for the boundary effect modification, k=1.3805×10-23 J/K is the Boltzmann constant, T is the absolute temperature of the fluid, μ is the viscosity, and dp is the particle diameter. Assuming that the properties of the solution, i.e., the temperature and viscosity, are constant, the diameters of the antibodycoated nanobeads will increase and the diffusion coefficient of the nanobeads will decrease when the antibody conjugates with the antigen. The diffusion coefficient of the nanobeads can be measured by diffusometry as mentioned in our previous work.28

Figure 7. Operation process of the preconcentrator and bubble-valves. (a) Microscopic view of the preconcentrator and bubble-valves. (b) Image of the preconcentrator. A depleted region began to form, and the fluorescent molecules began to accumulat.e (c) The preconcentrating plug completely formed when 25 V was applied at the two ends of the main channel. (d) and (e) The preconcentrating plug was moved into the microchamber by adjusting the applied voltages. (e) By turning on the piezoelectric transducer to excite the bubbles, the preconcentrating plug was captured inside the microchamber. (f) The fluorescent molecules slowly diffused to the entire microchamber.

Bead immunosensing analysis by diffusometry. In this section, we demonstrate application of the nanofluidic preconcentrator-integrated immunobead biosensor. Polystyrene beads (Merck™ XC030) with a diameter of 300 nm were used. The protocol, as mentioned in

Based on the Brownian motion theory, the probability of displacement is interpreted as mass diffusivity. The Brownian velocity is defined as the root mean square value of all observed nanobeads’ velocities. In our algorithm, every recognized nanobeads’ velocity was recorded and thus the Brownian velocity was statistically analyzed. For bead immunosensing, anti-CRP-coated nanobeads were mixed with samples and injected into the main channel to sense the antigen. Figure 8(a) shows measurement results of equilibrium Brownian velocities of anti-CRP-coated nanobeads when mixed with various concentrations of CRP solutions in the range of 0~8.4 μg/mL. From the experimental results, we found that the equilibrium Brownian velocity decreased as the sample concentration increased. This result indicates that there were more CRPs bound onto the anti-CRP-coated nanobeads, which resulted in that the nanobeads enlarging. Brownian velocities of the nanobeads with different CRP concentration solutions in an equilibrium state were plotted and linearly fitted in Fig. 8(b). In the linear-fit region, the equilibrium Brownian velocity of the nanobeads can be used to estimate the sample concentration.

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A negative control experiment using bare nanobeads in 8.4 µg/mL CRP was implemented to investigate whether or not the change of the solution viscosity due to the introduction of the CRP will affect the Brownian velocity. The averaged equilibrium Brownian velocity of the bare nanobeads in 8.4 µg/mL CRP solution is 6.23±0.11 µm/s, which is slightly lower than that the antibody-coated nanobeads in a solution without CRP.

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Samples with CRP concentrations of 105, 210, 420, and 840 pg/mL were mixed with anti-CRP-coated nanobeads and injected into the main channel. The equilibrium Brownian velocities of these anti-CRP-coated nanobeads were found to be indistinguishable from that without CRP as depicted in Figs. 8 and 9. The reason is that variations in the nanobead diameters were too small to be observed due to the low CRP concentration. After preconcentration, the Brownian velocities of the nanobeads in the four CRP samples decreased as shown in Fig. 9. These results indicated that a higher CRP concentration led to a moredramatic decrease in the Brownian velocity of the nanobeads. Moreover, the results revealed that the nanofluidic preconcentrator can be used to enhance the detection limit of the bead-based immunoassay to the pg/mL level by measuring the Brownian diffusion of the nanobeads. The averaged preconcentration factor of about 10,000 was obtained

Figure 9. Detection of low-abundance C-reactive protein (CRP) samples at four concentrations of 105, 210, 420, and 840 pg/mL, with and without preconcentration. The hollow symbols represent sample detection without preconcentration, while the solid symbols are detection results with preconcentration.

Conclusions

Figure 8. (a) Variations of the equilibrium Brownian velocity of the anti-C-reactive protein (CRP)-coated nanobeads with respect to different CRP concentrations. (b) Results of equilibrium Brownian velocities showing a linear decrease after reacting with different C-reactive protein (CRP) concentrations.

Bead immunosensing integrated with the preconcentrator.

A new biomarker detection method that integrates oscillating-bubble valves with a nanofluidic preconcentrator was successfully demonstrated. Acoustically excited oscillating bubble valves have several advantages, including simple fabrication with only one layer of a microfluidic channel, perfect channel-blocking effects, multiple valves which can be selectively turned on and off easily, and a high-pressure air source is not required, which has the potential for integration with portable devices. For the immunoassay, an ultra-low detection limit, of as low as 100

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pg/mL. The whole processes can be finished in 30 min including sample mixing and injection for 3 min, preconcentration for 10 min, valve operation for 2min, and sensing process for 15 min.

ACKNOWLEDGMENT We would like to thank Prof. Yen-Wen Lu’s helpful discussion concerning several technical problems. This work was supported by the Ministry of Science and Technology of Taiwan under grant numbers MOST 106-3114-E-002-003 and 106-2218-E-038-003. We would also like to thank the NEMS Research Center of National Taiwan University for facility support.

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A Novel Thermal Bubble Valves integrated Nanofluidic

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