Microfluidics plays an important role for the development of new sensor concepts aiming at the detection of target analytes in low-volume liquid samples.1,2 Microfluidic or Lab-On-a-Chip (LOC) devices have been applied over a wide range of chemical,3–5 biological6–9 and physical systems.10–12 Conventionally, single-layer microfluidic devices consist of two parts: one part comprises the structured channels and the other one resembles the cover or the lid. These two parts – one, containing all the functional fluidic channels and chambers, and the other, covering the exposed side and closing the channels – are produced separately and bonded afterwards, either permanently or non-permanently. The microfluidic counterpart sometimes bears some functionality. Sensors or other functional devices are usually processed on the planar cover (or bottom), which allows for standard fabrication processes (e.g. metal electrode deposition). In this case the sensing structures should remain intact during the bonding process. Additionally, the two structures often need to be accurately aligned to avoid obstruction of the sensor by the channel walls. There are many well-established strategies for permanent bonding of microfluidic devices,13–16 which can be roughly divided into two main categories: with and without using an additional adhesive interlayer. However, none of these techniques are universal and often depend on the nature of the materials, which are used to fabricate the channel structure and counter plate. Besides, if sensing elements are involved or the critical dimensions of the channels are on the lower micrometer scale, special limitations are imposed on the bonding process regarding materials, pressure, and temperature to avoid degradation or deformation of devices. Because of its popularity in the academic community, a big share of bonding strategies is dedicated to PDMS (polydimethylsiloxane) and PDMS/glass combinations. PDMS is the most widespread material for lab-based LOC devices. However, it has some major drawbacks such as chemical incompatibility with organic solvents, adsorption of hydrophobic molecules, short-term stability after surface treatment, water permeability17,18 and, also, challenges with scalability for mass production since molds are involved. Other, potentially more commercially viable, polymers have received less attention. Unlike for PDMS, for many of those materials bonding methods without the use of adhesives are either unavailable or can be damaging to the performance of the sensing elements. Therefore, in this case, an adhesive interlayer, which has good adhesion properties for both parts of the LOC device, is often applied for bonding. Among a variety of potential adhesives, UV-curable systems are quite popular in this respect,19–22 since they do not require high temperature or pressure for curing, enabling the use of polymers with low glass transition temperatures (Tg). However, it is often challenging to pattern such adhesives on small structures, whereas unpatterned adhesives may flow into and block the microfluidic channels. Overall, there are some important criteria that are generally required for an efficient bonding method, which include high bonding strength and bursting pressure, curing at room temperature and ambient pressure, resistivity to defects and dust particles (i.e. no requirement for dust-free cleanroom environment), alignment precision, rapid prototyping options, and easy handling. Rarely can all of these criteria be met by any single approach.
At the other frontier, printed electronics technologies, especially digital printing such as inkjet printing, have been recently gaining momentum. There is a plethora of publications on applications of functional inkjet printing in photovoltaics, displays, sensor development as well as printing of biological proteins, cells and tissues.23–27 The main advantages of inkjet printing are the facile design adaptability, relatively high lateral resolution and deposition of a large variety of materials from the liquid phase. Such features make inkjet printing a great tool for prototyping sensing platforms using a variety of functional inks. Moreover, the rise of affordable 3D-printing technologies such as stereolithography (SLA) has enabled the rapid prototyping of devices on a micrometer scale,28 although other rapid prototyping approaches have also been proposed.29 However, the resolution and surface roughness of printed structures are typically not sufficient to fabricate closed channels in a one-part device. Therefore, especially if one also plans to include a sensing element, efficient approaches for bonding to a polymer or glass substrate still have to be developed.
In this technical innovation, we demonstrate the use of inkjet printing for patterned deposition of UV-curable polymer inks for bonding of 3D-printed microfluidic devices and polymer foils. The same inks also serve as a dielectric passivation layer for conductive tracks in electrical or electrochemical sensors. Thus, passivation of printed sensors and deposition of adhesives can be achieved in one step. Having utilized inkjet printing in conjunction with SLA 3D-printing, we demonstrate a fully printed approach for rapid prototyping of microfluidic systems with electrical or electrochemical sensing capabilities using temperature-sensitive materials with low Tg.
Test bodies for bonding strength measurements were prepared by a desktop stereolithography (SLA/DLP) 3D printer (Miicraft, Hsinchu, Taiwan). All samples were designed in AUTOCAD 2013 (Autodesk Inc., USA) and converted into STL files. These structures were sliced in 2D layers using the 3D Miicraft printer software, which generates Portable Network Graphic images (PNG) to feed the DLP pico-projector (450 dpi). Samples were printed with 50 μm layer thickness using UV acrylate Clear Resin BV-003 (Young Optics Inc., Hsinchu, Taiwan) with a solid surface energy of 41 mN m−1 after curing.
The 3D-printed samples dimensions were 2.5 × 5 × 5 mm3 (H × W × L) with support on the top for connecting the weights for force measurement (see Fig. S1†). After printing, the samples were washed with ethanol to remove uncured resin, dried with nitrogen, and post-cured using a printer-integrated UV-lamp (18 W UVA lamp).
As substrates, we used polyethylene naphthalate (PEN) specially coated Optfine® PQA1M (Teijin DuPont Films) defect-free surface with a solid surface energy of 30 mN m−1, and Teonex® Q83 (Teijin DuPont Films) with a solid surface energy of 34 mN m−1. Prior to printing, Teonex® Q83 substrates were cleaned with ethanol and sonicated for 5 minutes in an ultrasonic bath, dried with nitrogen, and heated at 100 °C for 5 minutes. Due to the presence of protective foil, the PQA1M substrate was used as received. Both substrates were treated with oxygen plasma (30 W, 0.2 mbar for different time periods) (Nano, Diener Electronic GmbH).
An inkjet printer (OmniJet 300, UniJet Co., Republic of Korea) was used to print ink layers on the polymeric substrates. Two UV-curable inks were evaluated: one formulated in house, the other was a commercially available PA-1210 series (high resistivity, UV-curable ink, JNC Corporation, Tokyo, Japan). The UV-curable ink formulated in house was based on a PVP-co-PMMA polymer described elsewhere.30,31 The viscosity and the surface tension of the ink were measured with a viscometer (μVisc, RheoSense Inc., San Ramon, CA, US) and a tensiometer (Kino A3, USA Kino Industry Co., Shanghai, China), respectively. The viscosity and surface tension of the PVP-co-PMMA ink were adjusted to approximately 10 cps and 30 mN m−1 to comply with optimal jetting requirements. Before printing, the inks were filtered with a 0.45 μm PVDF syringe filter to prevent particles and gas bubbles reaching the cartridge. Dimatix DMC 10 pL cartridges were used for printing. The printing was usually done at a jetting frequency of 1 kHz and resolution in the range from 800 to 1700 dpi. For bonding strength measurements, squares of the same dimensions as the test bodies of the 3D samples were printed on the substrate, onto which immediately after finishing the printing, the 3D test bodies were placed. Next, to start the curing process, the samples were exposed to UV light (1.1 W cm−2) providing a good cross-linking polymerization and bonding the samples to the substrate. Prior to performing the adhesion tests, the bonded samples were left drying in ambient condition for 24 h to ensure complete evaporation of the ink solvent.
Homemade bonding strength and bursting pressure measurement systems were built to evaluate, respectively, the bonding and sealing quality between the 3D-printed sample and the substrate bonded with the UV-cured ink (see Fig. S2–S4†).
For demonstrating the microfluidic flow, a test block (3 × 15 × 10 mm3, H × W × L) with a built-in channel (0.7 × 0.8 × 15 mm3, H × W × L) was printed using the 3D printer. Afterwards, a designated structure replicating the outer dimensions of the microfluidic channel was printed using the PVP-co-PMMA ink onto the PQA1M substrate under the same conditions as mentioned before. After bonding, the fabricated device without any further surface modification was used to test the liquid flow and sealing, obtained using the presented rapid prototyping test process, by guiding whole blood flow under the effect of capillary forces only.
For passivation tests, conducting test structures were fabricated using a commercial silver nanoparticle ink (Silverjet DGP40LT-15C, Advanced Nano Products Co., Ltd) and a homemade carbon black formulation (based on Printex L6 from Grolman Group, Neuss, Germany) in water/glycol mixtures. Silver and carbon were printed on a PQA1M substrate as feedlines and microelectrodes, respectively. All inks were printed using 10 pL cartridges with a frequency of 1–2 kHz and sintered at 130 °C for 3 hours. Next, PVP-co-PMMA ink was printed as a dielectric material with a resolution of 1693 dpi, and subsequently UV-cured with a dosage of 1.1 W cm−2. Electrochemical experiments were carried out using a VSP-300 potentiostat, BioLogic Science Instruments. Cyclic voltammogram (CV) measurements were performed using 500 μM 1,1-ferrocene dimethanol, (Sigma-Aldrich) prepared in phosphate buffer saline (PBS) solution (pH 7.4). CV measurements were carried out by sweeping the electrode potential between −0.3 and 0.5 V vs. a Ag/AgCl reference electrode (Super Dri-ref SDR 2, World Precision Instruments) at a scan rate of 100 mV s−1.