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Solid-State Microfluidics with Integrated Thin-Film Acoustic Sensors Menglun Zhang, Jingze Huang, Yao Lu, Wei Pang, Hao Zhang, and Xuexin Duan ACS Sens., Just Accepted Manuscript • DOI: 10.1021/acssensors.8b00412 • Publication Date (Web): 24 Jul 2018 Downloaded from http://pubs.acs.org on July 26, 2018
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Solid-State Microfluidics with Integrated Thin-Film Acoustic Sensors Menglun Zhang,1,‡ Jingze Huang,2,‡ Yao Lu,1 Wei Pang,1,2,* Hao Zhang,2 and Xuexin Duan1,* 1
State Key Laboratory of Precision Measuring Technology and Instruments, Tianjin University, Tianjin 300072, China.
2
College of Precision Instrument and Opto-electronics Engineering, Tianjin University, Tianjin 300072, China.
Keywords: acoustic wave sensors; thin film microfluidics; monolithic integration; point-of-care; biosensors
Abstract: For point-of-care applications, integrating sensors into a microfluidic chip is a nontrivial task, since conventional detection modules are bulky and microfluidic chips are small in size, and their fabrication processes are not compatible. In this work, a solid-state microfluidic chip with on-chip acoustic sensors using standard thin-film technologies is introduced. The integrated chip is essentially a stack of thin films on silicon substrate, featuring compact size, electrical input (fluid control) and electrical output (sensor read-out). These features all contribute to portability. In addition, by virtue of processing discrete micro-droplets, the chip provides a solution to the performance degradation bottleneck of acoustic sensors in liquid-phase sensing. Label-free immunoassays in serum are carried out and the viability of the chip is further demonstrated by result comparison with commercial ELISA in prostate-specific antigen sensing experiments. The solidstate chip is believed to fit specific applications in personalized diagnostics and other relevant clinical settings where instrument portability matters.
One of the major objectives of microfluidics is to perform point-of-care (POC) diagnostic for personalized healthcare.[1] To replace traditional macroscale assays with a handheld instrument performing home or personal medical diagnosis, the instrument portability is a key factor. [2] Microfluidic technology takes the advantage of conducting complex assay protocols with reduced sample and reagent volume on a small chip. However, there are at least two aspects of constraint on practical use of microfluidics in POC applications. On one hand, the diagnostic output has to rely on detection components, which are usually in the form of optical microscope or other bulky imaging instruments. [3,4] On the other hand, conventional microfluidics uses auxiliary fluid pumps, which are not welcome in terms of over-all size and integration. Digital microfluidics (DMF) based on electrowetting-on-dielectric technology, a thin-film type device, has become a popular pump-free liquid handling technology in recent years. [5] Compared with the conventional microfluidics, it can electrically manipulate microdroplets with high control flexibility, which does not require extra valves and pumps. [6] Thus, it is a great benefit for developing POC tools. [7,8] Early investigations have developed DMF-based systems comprising off-chip detection modules, such as mass spectrometry[9] and chemiluminescence detection by photo multiplier tubes. [10] There is a trend to integrate on-chip sensors with DMF for better qualification of POC applications. [11-17] Thin-film piezoelectric acoustic sensors (e.g. thin-film bulk acoustic wave sensors, thin-film surface acoustic wave sensors, and thin-film Lamb wave sensors) have been extensively investigated in biological and chemical sensing for decades. [18-21] They feature small size, high sensitivity, and label-free detection. In addition, their signal acquisition and processing are essentially in electrical form, as electrical microsensors. These features make them ideal as integrated sensors for portable applications. Among them, thin-film bulk acoustic wave sensors have the highest sensitivity and smallest size. [22] In terms of compactness and integration, a solid-state microfluidic chip is proposed, referring to solid-state devices as being monolithic integrated circuits, fabricated using thin-film technologies. In this paper, thin-film bulk acoustic wave sensors are integrated into DMF as a stack of thin films on a silicon substrate, forming a so-called solid-state microfluidic chip. Like an integrated circuit chip, the input control and output sensor-readout of the solid-state microfluidic chip are all electrical signals. The chip is small, pump-free and optics-free for POC applications. The chip is demonstrated for label-free immunoassay with liquid-handling processed by DMF and target biomolecule detection by on-chip sensors.
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■ Concept and experimental Chip structure and working principle
Figure 1. (a) Schematic of the chip with DMF droplet manipulating path (silver), sensor array (pink), and tiny droplets containing a variety of biomolecular targets (black) and functionalization probes (red, green and blue). (b)-(c) Cross-sectional schematic illustrating the working principle. The manipulated droplet is off the sensor with acoustic energy trapping (b) and on top of the sensor with acoustic energy loss (c). (d) Photo of the fabricated chip with an area of 225 mm2. (e) Enlarged view of an onchip sensor with a size of 200 μm. (f) SEM image of the sensor cross section. RC and BR stand for resonating cavity and Bragg reflector, respectively. (g) Sensor signal of frequency shift by mass-loading effect. (h) Principle of high-Q sensing mode enabled by DMF.
Figure 1a shows the schematic of the integrated chip. Multiple tiny droplets on the order of μL can be manipulated (transporting, splitting, merging and mixing) by the DMF driving electrodes. On-chip microsensors are tightly embedded in the droplet driving pathways. Sample droplets are driven by DMF to sensors for target molecules detection. In addition, the sensor functionalization for the specific target detection can also be accomplished on chip by manipulating probe droplets. The chip is equipped with a sensor array, as shown in Figure 1a, so that sensors can be functionalized by many droplets containing different biomolecule probes. Therefore the chip is capable of conducting parallel and multiplexed detection, thanks to the DMF unique feature of individual droplet addressability. It should be noted that all the work described above are program-controlled on a small single chip. The chip working principle is illustrated in Figure 1b and Figure 1c. DMF manipulate droplets by controlling the surface tension with electric field known as electrowetting-on-dielectric. The electric voltage applied on the electrodes under the hydrophobic dielectric layer changes the local surface wettability from hydrophobicity to hydrophilicity and vice versa. Therefore, droplet movement will trace the sequentially addressed electrodes of DMF (i.e. manipulating path). Acoustic wave sensors at the detection sites are surrounded by the DMF electrodes. High frequency (~GHz) electrical signal is
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applied to the piezoelectric thin film sandwiched between the top and bottom electrodes of the sensors, generating mechanical vibrations which amplitude is about 1 angstrom. [23] Large biomolecules absorbed on the sensor surface disturb the infinitesimal vibration. Hence a thin-film bulk acoustic wave sensor is sensitive to the absorbed mass, reaching a mass sensitivity of several hundreds of cm2g-1, compared to that of several tens of cm2g-1 for a quartz crystal microbalance (QCM). [18] The coupling between electrical and mechanical energy modulates the electrical input impedance of the sensor and gives rise to resonance (peak and valley) on frequency spectrum. As shown in Figure 1g, absorbed biomolecules are detected by the resonant frequency shift. Classic Sauerbrey equation shows that small resonant frequency shift (< 2%) is proportional to the added mass on the resonant sensor. [24] However, high level of sensor noise will hinder perception of small signals generated by biomolecule absorption. The noise is inversely proportional to the sensor qualify factor Q, which is determined by its energy trapping ability. When the sensor resonates in air (no droplet contact), the acoustic energy is trapped in the sensor (Figure 1b), and hence a high Q and low level of noise. When the sensor is in contact with a droplet (i.e. on-incubation), large amount of acoustic energy leak into the liquid and get lost (Fib. 1c), and hence a low Q and degraded noise. [25,26] To facilitate both biomolecular binding and low noise level sensing, the chip adopts a high-Q sensing mode: the sensing signal is acquired after the incubation process. When the incubation droplet is driven off the sensor, as shown in Figure 1h, the sensor resonates in air again and recovers its high Q to the pre-incubation state. The difference between the pre-incubation sensor and the post-incubation sensor is that the desired biomolecules are bound on the sensor surface. Therefore, the sensor operating in the high-Q mode is able to detect biomolecule binding. In fact, the high-Q mode of the acoustic sensor is enabled by the flexible droplet manipulation ability, which is uniquely endowed by the digital feature of DMF and nontrivial to achieve with continuous microfluidics. It should be noted that inputs and outputs of the chip are all electrical signals, as shown in Figure 1b. The DMF is controlled by a voltage source and addressing circuit. Although the sensor is an acoustic device, the acoustic energy is powered by and transformed back to the electrical form via piezoelectric thin film. The electrical signal of resonant frequency shift is acquired by a network analyser. Therefore, it has a great potential to shrink down the auxiliary components (i.e. DMF control and sensor read-out) into a more compact size, which is an essential feature to point-of-care applications. A chip demonstrating the above concept is manufactured. Figure 1d shows the photo of the chip. The chip has an electrode-defined active size of 15 mm and a thickness of 400 μm. DMF electrodes are electrically connected to the voltage control board, which is programmed to control droplet manipulation. There are four droplet pathways with three sensors embedded. Figure 1e shows more details about the acoustic microsensor. The sensor has square Bragg reflector with 200 μm in length. The pink area within Bragg reflector is the sensing area with a length of ~100 μm, where the structure is resonating at 2.5 GHz and sensitive to biomolecular binding. In thickness dimension, the sensor is composed of a stack of thin films (total ~5μm) on the silicon substrate (so-called solid-state acoustic wave sensor), as shown in Figure 1f. Bragg reflector uses alternatively stacked high and low acoustic impedance films on the substrate and each layer thickness is roughly a quarter of the acoustic wavelength in respective film at 2.5 GHz. Bottom electrode, piezoelectric thin film, top electrode and functionalization film are sequentially stacked to form acoustic resonating cavity on Bragg reflector. Bragg reflector reflects the acoustic energy back into the acoustic resonating cavity, preventing the energy leaking into silicon substrate. The solid-state acoustic sensor is different from the air cavity acoustic sensors used in our previous work. [27] The earlier demonstrated chip is not suitable for liquid-phase sensing, because liquid will occasionally leak into the air cavity through release holes, leading to random probe functionalization and unstable target signal. On the contrary, the chip with solid-state acoustic sensors gets rid of the problem by replacing the air cavity with solid Bragg reflector thin films, enabling stable biomolecule sensing.
Device fabrication The integrated chip fabrication uses standard thin-film fabrication process, as shown in Figure 2. Alternative thin films of SiO2/Mo/SiO2/Mo/SiO2 are deposited on a silicon substrate as the Bragg reflector (Figure 2a). The thickness of each layer is 650 nm. Then a 200 nm Mo film is sputtering deposited and patterned to form the actuation electrodes of DMF and the bottom electrode of the sensor on the Bragg reflector (Figure 2b). A uniform 1.1 μm AlN film is sputtering deposited on the electrodes to serve as both the piezoelectric layer of sensors and the dielectric layer of DMF. Afterwards, 200 nm Mo and 200 nm AlN films are sequentially deposited and patterned to form the top electrode and functionalization film of the sensor (Figure 2c). Then a 100 nm SiO2 film is deposited and patterned to solely expose the sensing area of the sensor. Finally, a 30 nm Teflon film is spin-coated and patterned by lift-off process to expose the sensing area (Figure 2d). Since photoresist developer used in the Teflon lift-off process etches AlN film, the thin SiO2 film is served as a barrier layer to keep the AlN film intact from the photoresist developer. The chip is ready for further chemical modification, biological functionalization and subsequent sensing procedures.
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Figure 2. Fabrication process of the solid-state chip. (a) Bragg reflector deposition; (b) bottom electrodes preparation; (c) preparation of piezoelectric layer, top electrode and functionalization layer; (d) hydrophobic Teflon thin film patterning.
Label-free and multiplexed IgG sensing In the experiment, different probe IgG antibodies are first linked onto the sensor array, and then target IgG antibodies in serum samples are detected for label-free and multiplexed sensing. The chip surface is cleaned by acetone, isopropyl alcohol, and deionized (DI) water sequentially before functionalization. Then the sensor surface is further cleaned with piranha solution within 30s. (3-Aminopropyl)triethoxysilane (APTES, Aladdin) is evaporated onto the surface by gas phase deposition in a heated vacuum chamber (labKote) for 15 mins (125℃, 5.7 Torr). Transformation of amino-terminated layer to an isothiocynated-bearing layer is accomplished by exposure to 1,4-phenylene diisothiocyanate (PDC, J&K Scientific) solution in ethanol at 65℃ for 1h. Afterwards, the chip surface is rinsed manually with copious amounts of ethanol, DI water sequentially, and dried in compressed nitrogen gas. After the chemical modification, biological specific functionalization is conducted. For the functionalization of IgG (Bio-synthesis Biotechnology) detection experiment, three droplets containing different probe antibodies (goat antimouse IgG, goat anti-rabbit IgG, goat anti-human IgG) are driven onto three sensors, respectively. The antibody droplets are incubated on the sensing areas of the sensor array for 30 mins. After incubation, the antibody droplets are driven away and the sensor surfaces are rinsed automatically. More specifically, by automatically moving DI water droplets on and off the sensor surface several times and repeating this step with fresh DI water droplets, the unbound molecules are rinsed off the sensor surface. After the specific functionalization of sensors with anti-IgGs, serum sample containing FITC labelled rabbit IgG as a target is detected by the chip. In detail, FITC-rabbit IgG is dissolved in 10% fetal bovine serum (FBS, Gibco). A 5 μg/ml FITC-rabbit IgG droplet is first driven onto the sensor array for an incubation of 20 mins, followed by automatic DI water droplet rinsing. Then a droplet of 100 μg/ml anti-rabbit IgG is driven onto the sensor array and allowed for an incubation of 20 mins, followed by another round of automatic DI water droplet rinsing. PSA ELISA experiment To further demonstrate our chip, a different set of experiments are carried out by comparing our PSA detection results with commercial fluorescence absorbance method. For the functionalization of PSA (Linc-Bio), three droplets containing 200 μg/ml anti-PSA are driven onto the respective sensors. The remaining steps are the same with the above IgG functionalization. The enzyme-linked immunosorbent assay (ELISA) is conducted as following. First, the sensor surface is immobilized with anti-PSA probe as the antibody functionalization. Then three droplets of different concentrations (0 μg/ml, 5 μg/ml, 20 μg/ml) of PSA droplets are driven onto the sensor array and incubated for 45 mins at 37℃. After automatic DI water droplet rinsing, three droplets containing horseradish peroxidase (HRP) conjugated anti-PSA (HRP-anti-
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PSA) are driven onto the sensor array and incubated for 45 mins at 37℃. After another round of DI water droplet rinsing, three droplets containing chromogenic TMB substrate are driven onto the sensors and incubated for 10 mins at 37℃. At last, the TMB substrate droplet merges with a H2SO4 droplet (2 M) and the merged droplet is added immediately into three wells of a 96-well microplate. The absorbance at 450 nm of each well is obtained by a microplate reader (Multiskan GO, Thermo Scientific).
Sensor signal recording As shown in Figure 1g, the biomolecules absorbed on the acoustic sensor surface cause resonant frequency shift by mass-loading effect. The extent of frequency shift reflects the concentration of the target biomolecules in the liquid droplets. In the multi-functionalization step, the first resonant frequency is recorded by the network analyzer (Agilent E5061B) before the biomolecule functionalization is carried out. After probe incubation and washing off the unbound molecules, the second resonant frequency of the same sensor is recorded again. The difference between the two frequencies (i.e. frequency shift) is the signal proportional to the mass of the probe molecules absorbed on the sensor surface. The signals of the sensor array with different functionalization are recorded in a parallel way. In the target detection step, similar to the multi-functionalization step, the first resonant frequency is recorded just before the sample droplet incubation. After the target incubation and washing off the unbound molecules, the second resonant frequency of the same sensor is recorded again. The frequency shift is the response of the target molecules. In the FITC-rabbit IgG detection experiment, the FITC fluorescence readout is used as a comparison to the frequency shift signal. The fluorescence image is captured by a fluorescence microscope (Olympus BX 53 with DP73 digital camera). DMF operation To manipulate the droplets, AC voltage (40~70 VRMS, 1 kHz) is applied to the DMF electrodes. All protein solutions used in the experiments are added with Pluronic F127 (Fanbo Biochemicals) to facilitate smooth droplet manipulation. In addition, a small sealed room with 100% relative humidity is created surrounding the chip, which eliminates the evaporation problem of the tiny droplets. The cover of the room is made transparent for droplet manipulation monitoring and fluorescence image capture.
■ Results and discussion Chip working demonstration In order to show the microfluidic processes during work clearly, the droplet working flow is demonstrated on the chip, as shown in Figure 3. Three functionalization droplets and a sample droplet are manually loaded at the loading sites (Figure 3a). Then all the remaining procedure is undertaken by DMF automatically. The functionalization droplets are driven to the sensor array for probe binding in parallel. The droplets are positioned precisely without touching each other (Figure 3b). It guarantees a high-density sensor array used in multiplexed target detections where sensors are functionalized with different probes. After the probe incubation, the functionalization droplets are driven back to their loading sites and the sample droplet containing target molecules is driven onto the sensor array (Figure 3c). Since the sample droplet covers all the sensors and the target molecule binding process takes place in parallel, the sample volume and detection time are substantially decreased. After the target incubation, the sample droplet is driven back to its loading site. Although not shown here, it should be noted that in the following practical detection experiments, DI water rinsing is dispensable and carried out automatically by the chip after each binding step, e.g. probe binding or target binding. In addition, the sample droplet is slightly larger than the functionalization droplets. In a multiplexed detection, the functionalization droplets have to be small enough to get rid of touching each other in the probe binding step, as shown in Figure 3b; on the other hand, the sample droplet has to be large enough to cover the three sensors simultaneously, as shown in Figure 3c. Therefore, the width of DMF functionalization pathways is designed to be larger than that of sample deliver pathway, as shown in Figure 1d. The high-Q mode is demonstrated by assessing quality factor Q and corresponding sensor signal noise when the sensor is loaded and unloaded with a DI water droplet. As shown in Figure 3d, the sensor shows a high Q of ~2000; the resonant frequency barely changes with time (standard deviation=0.02 MHz). Right after the droplet is loaded on the sensor, Q dramatically drops to ~30; the resonant frequency fluctuates wildly with time (standard deviation=0.3 MHz), which indicates a higher noise level compared to that before droplet loading. When the droplet is driven off the sensor, Q and signal noise quickly recover back to ~2000 and 0.02 MHz level, respectively. The chip uses the high Q nature of the resonant sensor in air and a fast air-liquid environment transition of DMF to guarantee a high sensing precision. The sensor shows a frequency shift of ~5 MHz when the droplet is loaded, because the mass of the droplet changes the resonant frequency of the sensor.
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Figure 3. Demonstration of the chip working. (a) All droplets are at their loading sites. (b) Three functionalization droplets (~10 μL) containing different probe biomolecules are driven onto the sensor array for probe binding in parallel. (c) A relatively larger sample droplet (~20 μL) is driven onto the sensor array for target biomolecules binding. (d) Real-time changes of quality factor Q and frequency shift as a DI water droplet is loaded on and later driven off a sensor.
Label-free sensing in serum The chip can be used in label-free and multiplexed biomolecule detection. As a proof of concept, antibody detections in serum sample are carried out. The sensor array on a single chip is first functionalized with different probes. Sensor A, B and C in the array is immobilized with anti-rabbit IgG, anti-mouse IgG and anti-human IgG, respectively. Then the serum sample containing the target biomolecules (rabbit IgG) is driven onto the sensor array. Since rabbit IgG specifically binds to anti-rabbit IgG, the target molecules will only bind on sensor A basically. After that, a reporting droplet containing anti-rabbit IgG is then driven onto the sensor array. Finally, an “anti IgG-IgG-anti IgG” sandwich structure is formed on the sensor surface. The signals of the three sensors correspond to every binding event in the multiplexed detection experiment, where the nature of biomolecular specific binding plays an important role.
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Figure 4. Experimental results of functionalization, target binding and reporter binding. (a) Sensor signals of frequency shift and the corresponding schematic after functionalization. (b) Fluorescent images, frequency shifts, and the corresponding schematic after target biomolecule binding. (c) Frequency shifts and the corresponding schematic after reporting biomolecule binding. The colour of green, blue and red represent rabbit, mouse and human IgGs and anti-IgGs, respectively. Error bars are calculated from standard deviations of repeated experimental results.
Figure 4a shows the results in the functionalization step. After functionalization and rinsing, the frequency shift of each sensor indicates a successful immobilization of the probe antibody on the sensor surface. The frequency shifts are all around 500 kHz for the three sensors, which means that the surface concentrations of the three immobilized probes are substantially equal. The equal surface concentration of different probes is important in the multi-functionalization step; otherwise it will cause a confusing interpretation of the sensor signal in the target detection step. In the unequal case, a large signal of a sensor will result from either a specific binding of the target to the specific probe, or a higher surface concentration of the probe which may lead to a higher level of non-specific binding. In the target binding step, serum sample containing FITC labelled rabbit IgG is introduced. Figure 4b shows that the frequency shifts of the three sensors are similar (~170 kHz). It indicates that the frequency signals could not resolve the target binding, since interference biomolecules (e.g. BSA and bovine IgG) present in the FBS sample and rabbit IgG will bind non-specifically to probes of anti-mouse IgG and anti-human IgG, besides specific binding of anti-rabbit IgG and rabbit IgG. However, with fluorescent FITC label, we confirm that target molecules of rabbit IgG prefer to bind on sensor A, as shown in Figure 4b. Much lower fluorescent intensity of sensor B and sensor C indicates little non-specific binding of rabbit IgG. It should be noted that, in practical detection applications, the target is only resolved by frequency shift of the reporter. Fluorescent label used here is merely used to demonstrate the correct binding. Finally, the reporter of anti-rabbit IgG is introduced. By detecting the binding of top anti-rabbit IgG, the target of rabbit IgG is quantified. In Fig 4c, sensor A shows a large frequency shift (~90 kHz), while sensor B and sensor C show small frequency shifts (
