Complex three-dimensional high aspect ratio microfluidic network manufactured in combined PerMX dry-resist and SU-8 technology

In this paper we present a new fabrication method that combines for the first time popular SU-8 technology and PerMX dry-photoresist lamination for the manufacturing of high aspect ratio three-dimensional multi-level microfluidic networks. The potential of this approach, which further benefits from wafer-level manufacturing and accurate alignment of fluidic levels, is demonstrated by a highly integrated three-level microfluidic chip. The hereby achieved network complexity, including 24 fluidic vias and 16 crossing points of three individual microchannels on less than 13 mm(2) chip area, is unique for SU-8 based fluidic networks. We further report on excellent process compatibility between SU-8 and PerMX dry-photoresist which results in high interlayer adhesion strength. The tight pressure sealing of a fluidic channel (0.5 MPa for 1 h) is demonstrated for 150 μm narrow SU-8/PerMX bonding interfaces.     

Focusing-enhanced mixing in microfluidic channels

A focusing-based microfluidic mixer was studied. The micromixer utilizes the focusing process required for cytometry to reduce the diffusion distance of molecules to be mixed in order to facilitate the passive diffusion-controlled mixing process. It was found that both the high flow rate ratio of the sheath flow to the flows to be mixed and the low flow rate of the mixing fluids resulted in the short mixing length required within the microfluidic channel. It was shown that a complete mixing was achieved within a distance of 4 mm in the micromixer for the focused mixing fluids at a flow rate of 2 μl∕min and a flow rate ratio of the sheath flow to the flows to be mixed at 4:1. The mixer described here is simple and can be easily fabricated and controlled.     

Surface patterning of bonded microfluidic channels

Microfluidic channels in which multiple chemical and biological processes can be integrated into a single chip have provided a suitable platform for high throughput screening, chemical synthesis, detection, and alike. These microchips generally exhibit a homogeneous surface chemistry, which limits their functionality. Localized surface modification of microchannels can be challenging due to the nonplanar geometries involved. However, chip bonding remains the main hurdle, with many methods involving thermal or plasma treatment that, in most cases, neutralizes the desired chemical functionality. Postbonding modification of microchannels is subject to many limitations, some of which have been recently overcome. Novel techniques include solution-based modification using laminar or capillary flow, while conventional techniques such as photolithography remain popular. Nonetheless, new methods, including localized microplasma treatment, are emerging as effective postbonding alternatives. This Review focuses on postbonding methods for surface patterning of microchannels.     

Study on surface properties of PDMS microfluidic chips treated with albumin

Electrokinetic properties and morphology of PDMS microfluidic chips intended for bioassays are studied. The chips are fabricated by a casting method followed by polymerization bonding. Microchannels are coated with 1% solution of bovine serum albumin (BSA) in Tris buffer. Albumin passively adsorbs on the PDMS surface. Electrokinetic characteristics (electro-osmotic velocity, electro-osmotic mobility, and zeta potential) of the coated PDMS channels are experimentally determined as functions of the electric field strength and the characteristic electrolyte concentration. Atomic force microscopy (AFM) analysis of the surface reveals a “peak and ridge” structure of the protein layer and an imperfect substrate coating. On the basis of the AFM observation, several topologies of the BSA-PDMS surface are proposed. A nonslip mathematical model of the electro-osmotic flow is then numerically analyzed. It is found that the electrokinetic characteristics computed for a channel with the homogeneous distribution of a fixed electric charge do not fit the experimental data. Heterogeneous distribution of the fixed electric charge and the surface roughness is thus taken into account. When a flat PDMS surface with electric charge heterogeneities is considered, the numerical results are in very good agreement with our experimental data. An optimization analysis finally allowed the determination of the surface concentration of the electric charge and the degree of the PDMS surface coating. The obtained findings can be important for correct prediction and possibly for robust control of behavior of electrically driven PDMS microfluidic chips. The proposed method of the electro-osmotic flow analysis at surfaces with a heterogeneous distribution of the surface electric charge can also be exploited in the interpretation of experimental studies dealing with protein-solid phase interactions or substrate coatings.     

Transport of particles and microorganisms in microfluidic channels using rectified ac electro-osmotic flow

A new method is demonstrated to transport particles, cells, and other microorganisms using rectified ac electro-osmotic flows in open microchannels. The rectified flow is obtained by synchronous zeta potential modulation with the driving potential in the microchannel. Experiments were conducted to transport both neutral, charged particles, and microorganisms of various sizes. A maximum speed of 50 μm∕s was obtained for 8 μm polystyrene beads, without any electrolysis, using a symmetrical square waveform driving electric field of 5 V∕mm at 10 Hz and a 360 V gate potential with its polarity synchronized with the driving potential (phase lag=0°).     

Hydrodynamic resistance and mobility of deformable objects in microfluidic channels

This work reports experimental and theoretical studies of hydrodynamic behaviour of deformable objects such as droplets and cells in a microchannel. Effects of mechanical properties including size and viscosity of these objects on their deformability, mobility, and induced hydrodynamic resistance are investigated. The experimental results revealed that the deformability of droplets, which is quantified in terms of deformability index (D.I.), depends on the droplet-to-channel size ratio ρ and droplet-to-medium viscosity ratio λ. Using a large set of experimental data, for the first time, we provide a mathematical formula that correlates induced hydrodynamic resistance of a single droplet ΔR_d with the droplet size ρ and viscosity λ. A simple theoretical model is developed to obtain closed form expressions for droplet mobility ? and ΔR_d. The predictions of the theoretical model successfully confront the experimental results in terms of the droplet mobility ? and induced hydrodynamic resistance ΔR_d. Numerical simulations are carried out using volume-of-fluid model to predict droplet generation and deformation of droplets of different size ratio ρ and viscosity ratio λ, which compare well with that obtained from the experiments. In a novel effort, we performed experiments to measure the bulk induced hydrodynamic resistance ΔR of different biological cells (yeast, L6, and HEK 293). The results reveal that the bulk induced hydrodynamic resistance ΔR is related to the cell concentration and apparent viscosity of the cells.     

Low cost fabrication and assembly process for re-usable 3D polydimethylsiloxane (PDMS) microfluidic networks

A method to easily manufacture and assemble a polydimethylsiloxane (PDMS) based microfluidic device is described. The method uses low cost materials and re-usable laser cut polymethyl methacrylate (PMMA) parts. In addition, the thickness of PDMS layers can be controlled and both PDMS layer surfaces are flat, which allows for multi-layer PDMS structures to be assembled. The use of mechanical clamping to seal the structure allows for easy cleaning and re-use of the manufactured part as it can be taken apart at any time. In this way, selected layers can be re-used or replaced. The process described can be easily adopted and utilised without the need for any costly clean room facilities or equipment such as oxygen bonders, making it ideal for laboratories, universities, and classrooms exploring microfluidics applications.     

Microwave frequency sensor for detection of biological cells in microfluidic channels

We present details of an apparatus for capacitive detection of biomaterials in microfluidic channels operating at microwave frequencies where dielectric effects due to interfacial polarization are minimal. A circuit model is presented, which can be used to adapt this detection system for use in other microfluidic applications and to identify ones where it would not be suitable. The detection system is based on a microwave coupled transmission line resonator integrated into an interferometer. At 1.5 GHz the system is capable of detecting changes in capacitance of 650 zF with a 50 Hz bandwidth. This system is well suited to the detection of biomaterials in a variety of suspending fluids, including phosphate-buffered saline. Applications involving both model particles (polystyrene microspheres) and living cells—baker’s yeast (Saccharomyces cerevisiae) and Chinese hamster ovary cells—are presented.     

On-chip microfluidic biosensor using superparamagnetic microparticles

In this paper, an integrated solution towards an on-chip microfluidic biosensor using the magnetically induced motion of functionalized superparamagnetic microparticles (SMPs) is presented. The concept of the proposed method is that the induced velocity on SMPs in suspension, while imposed to a magnetic field gradient, is inversely proportional to their volume. Specifically, a velocity variation of suspended functionalized SMPs inside a detection microchannel with respect to a reference velocity, specified in a parallel reference microchannel, indicates an increase in their non-magnetic volume. This volumetric increase of the SMPs is caused by the binding of organic compounds (e.g., biomolecules) to their functionalized surface. The new compounds with the increased non-magnetic volume are called loaded SMPs (LSMPs). The magnetic force required for the manipulation of the SMPs and LSMPs is produced by current currying conducting microstructures, driven by a programmable microcontroller. Experiments were carried out as a proof of concept. A promising decrease in the velocity of the LSMPs in comparison to that of the SMPs was measured. Thus, it is the velocity variation which determines the presence of the organic compounds in the sample fluid.     

Fabrication of microfluidic devices using polydimethylsiloxane

Polydimethylsiloxane (PDMS) is nearly ubiquitous in microfluidic devices, being easy to work with, economical, and transparent. A detailed protocol is provided here for using PDMS in the fabrication of microfluidic devices to aid those interested in using the material in their work, with information on the many potential ways the material may be used for novel devices.     

Development of vertical SU-8 microtubes integrated with dissolvable tips for transdermal drug delivery

Polymer-based microneedles have drawn much attention in the transdermal drug delivery resulting from their flexibility and biocompatibility. Traditional fabrication approach deploys various kinds of molds to create sharp tips at the end of needles for the penetration purpose. This approach is usually time-consuming and expensive. In this study, we developed an innovative fabrication process to make biocompatible SU-8 microtubes integrated with biodissolvable maltose tips as novel microneedles for the transdermal drug delivery applications. These microneedles can easily penetrate the skin''s outer barrier represented by the stratum corneum (SC) layer. The drug delivery device of mironeedles array with 1000 μm spacing between adjacent microneedles is proven to be able to penetrate porcine cadaver skins successfully. The maximum loading force on the individual microneedle can be as large as 7.36 ± 0.48N. After 9 min of the penetration, all the maltose tips are dissolved in the tissue. Drugs can be further delivered via these open biocompatible SU-8 microtubes in a continuous flow manner. The permeation patterns caused by the solution containing Rhodamine 110 at different depths from skin surface were characterized via a confocal microscope. It shows successful implementation of the microneedle function for fabricated devices.     

Tunable visual color filter using microfluidic grating

This paper reports a tunable visual color filter based on a microfluidic transmission grating. The grating lines are formed by the microflows in an array of evenly spaced straight microchannels. In experimental study, the transmission of white light measures a shift of visual color from red to blue in the zeroth order diffraction in response to a change of the refractive index from 1.3290 to 1.3782 in the microflows. The merit of large tunability of transmission peak (Δλ=408 nm) makes this grating potential for various applications in biological and chemical measurements, such as space- and time-resolving micropattern spectrophotometers and separation of the fluorescence from the excitation.     

Development of vertical SU-8 microneedles for transdermal drug delivery by double drawing lithography technology

Polymer-based microneedles have drawn much attention in transdermal drug delivery resulting from their flexibility and biocompatibility. Traditional fabrication approaches are usually time-consuming and expensive. In this study, we developed a new double drawing lithography technology to make biocompatible SU-8 microneedles for transdermal drug delivery applications. These microneedles are strong enough to stand force from both vertical direction and planar direction during penetration. They can be used to penetrate into the skin easily and deliver drugs to the tissues under it. By controlling the delivery speed lower than 2 μl/min per single microneedle, the delivery rate can be as high as 71%.Microelectromechanical systems (MEMS) technology has enabled wide range of biomedical devices applications, such as micropatterning of substrates and cells,¹ microfluidics,² molecular biology on chips,³ cells on chips,⁴ tissue microengineering,⁵ and implantable microdevices.⁶ Transdermal drug delivery using MEMS based devices can delivery insoluble, unstable, or unavailable therapeutic compounds to reduce the amount of those compounds used and to localize the delivery of potent compounds.⁷ Microneedles for transdermal drug delivery are increasingly becoming popular due to their minimally invasive procedure,⁸ promising chance for self-administration,⁹ and low injury risks.¹⁰ Moreover, since pharmaceutical and therapeutic agents can be easily transported into the body through the skin by microneedles,^{11, 12} the microneedles are promising to replace traditional hypodermic needles in the future. Previously, various microneedles devices for transdermal drug delivery applications have been reported. They have been successfully fabricated by different materials, including silicon,¹³ stainless steel,¹⁴ titanium,¹⁵ tantalum,¹⁶ and nickel.¹⁷ Although microneedles with these kinds of materials can be easily fabricated into sharp shape and offer the required mechanical strength for penetration purpose, such microneedles are prone to be damaged¹⁸ and may not be biocompatible.¹⁹ As a result, polymer based microneedles, such as SU-8,^{20, 21} polymethyl meth-acrylate (PMMA),^{22, 23} polycarbonates (PCs),^{24, 25} maltose,^{26, 27} and polylactic acid (PLA),^{28, 29} have caught more and more attentions in the past few years. However, in order to obtain ultra-sharp tips for penetrating the barrier layer of stratum corneum,³⁰ conventional fabrication technologies, for instances, PDMS (Polydimethylsiloxane) molding technology,^{31, 32} stainless steel molding technology,³³ reactive ion etching technology,³⁴ inclined UV (Ultraviolet) exposure technology,³⁵ and backside exposure with integrated lens technology³⁶ are time-consuming and expensive. In this paper, we report an innovative double drawing lithography technology for scalable, reproducible, and inexpensive microneedle devices. Drawing lithography technology³⁷ was first developed by Lee et al. They leveraged the polymers'' different viscosities under different temperatures to pattern 3D structures. However, it required that the drawing frames need to be regular cylinders, which is not proper for our devices. To solve the problem, the new double drawing lithography is developed to create sharp SU-8 tips on the top of four SU-8 pillars for penetration purpose. Drugs can flow through the sidewall gaps between the pillars and enter into the tissues under the skin surface. The experiment results indicate that the new device can have larger than 1N planar buckling force and be easily penetrated into skin for drugs delivery purpose. By delivering glucose solution inside the hydrogel, the delivering rate of the microneedles can be as high as 71% when the single microneedle delivery speed is lower than 2 μl/min.An array of 3 × 3 SU-8 supporting structures was patterned on a 140 μm thick, 6 mm × 6 mm SU-8 membrane (Fig. ​(Fig.1a).1a). Each SU-8 supporting structure included four SU-8 pillars and was 350 μm high. The four pillars were patterned into a tubelike shape on the membrane (Fig. ​(Fig.1b).1b). The inner diameter of the tube was 150 μm, while the outer diameter was 300 μm. SU-8 needles of 700 μm height were created on the top of SU-8 supporting structures to ensure the ability of transdermal penetration. Two PDMS layers were bonded with SU-8 membrane to form a sealed chamber for storing drugs from the connection tube. Once the microneedles entered into the tissue, drugs could be delivered into the body through the sidewall gaps between the pillars (Fig. ​(Fig.1c1c).Open in a separate windowFigure 1Schematic illustration of the SU-8 microneedles. (a) Overview of the whole device; (b) SU-8 supporting structures made of 4 SU-8 pillars; and (c) enlarged view of a single SU-8 microneedle.The fabrication process of SU-8 microneedles is shown in Fig. ​Fig.2.2. SU-8 microneedles fabrication started from a layer of Polyethylene Terephthalate (PET, 3M, USA) film pasted on the Si substrate by sticking the edge area with kapton tape (Fig. ​(Fig.2a).2a). The PET film, a kind of transparent film with poor adhesion to SU-8, was used as a sacrificial layer to dry release the final device from Si substrate. A 140 μm thick SU-8 layer was deposited on the top of this PET film. To ensure a uniform surface of this thick SU-8 layer, the SU-8 deposition was conducted in two steps coating. After exposed under 450 mJ/cm² UV, the membrane pattern could be defined (Fig. ​(Fig.2b).2b). In order to ensure an even surface for following spinning process, another 350 μm SU-8 layer was directly deposited on this layer in two steps without development. With careful alignment, an exposure of 650 mJ/cm² UV energy was performed on this 350 μm SU-8 layer to define the SU-8 supporting structures (Fig. ​(Fig.2c).2c). The SU-8 structure could be easily released from the PET substrate by removing the kapton tape and slightly bending the PET film. Two PDMS layers were bonded with this SU-8 structure by a method reported by Zhang et al.³⁸ (Fig. ​(Fig.2d2d).Open in a separate windowFigure 2Fabrication process for SU-8 microtubes. (a) Attaching a PET film on the Si substrate; (b) exposing the first layer of SU-8 membrane without development; (c) depositing and patterning two continuous SU-8 layers as sidewall pillars; (d) releasing the SU-8 structure from the substrate and bonding it with PDMS; (e) drawing hollowed microneedles on the top of supporting structures; (f) baking and melting the hollowed microneedles to allow the SU-8 flow in the gaps between pillars; and (g) drawing second time on the top of the melted SU-8 flat surface to get microneedles.In our previous work,³⁹ we used one time stepwise controlled drawing lithography technology for the sharp tips integration. However, since the frame used to conduct drawing process in present study is a four-pillars structure rather than a microtube, the conventional drawing process can only make a hollowed tip but not a solid tip structure (Fig. ​(Fig.3).3). This kind of tip was fragile and could not penetrate skin in the practical testing process. To solve the problem, we developed an innovative double drawing lithography process. After bonding released SU-8 structure with PDMS layers (Fig. ​(Fig.2d),2d), we used it to conduct first time stepwise controlled drawing lithography³⁷ and got hollowed tips (Fig. ​(Fig.2e).2e). Briefly, the SU-8 was spun on the Si substrate and kept at 95 °C until the water inside completely vaporized. Device of SU-8 supporting structures was fixed on a precision stage. Then, the SU-8 supporting structures were immersed into the SU-8 by adjusting the precision state. The SU-8 were coated on the pillars'' surface. Then, the SU-8 supporting structures were drawn away from the interface of the liquid maltose and air. After that, the temperature and drawing speed were increased. Since the SU-8 was less viscous at higher temperature, the connection between the SU-8 supporting structures and surface of the liquid SU-8 became individual SU-8 bridge, shrank, and then broke. The end of the shrunk SU-8 bridge forms a sharp tip on the top of each SU-8 supporting structure when the connection was separated. After the hollowed tips were formed in the first step drawing process, the whole device was baked on the hotplate to melt the hollowed SU-8 tips. Melted SU-8 reflowed into the gaps between four pillars and the tips became domes (Fig. ​(Fig.2f).2f). Then, a second drawing process was conducted on the top of melted SU-8 to form sharp and solid tips (Fig. ​(Fig.2g).2g). The final fabricated device is shown in Fig. ​Fig.44.Open in a separate windowFigure 3A hollowed SU-8 microneedle fabricated by single drawing lithography technology (scale bar is 100 μm).Open in a separate windowFigure 4Optical images for the finished SU-8 microneedles.During the double drawing process, as long as the heated time and temperature were controlled, the SU-8 flow-in speed of SU-8 inside the gaps could be precisely determined. The relationship between baking temperature and flow-in speed was studied. As shown in Fig. ​Fig.5,5, the flow-in speed is positive related to the baking temperature. The explanation for this phenomena is that the SU-8''s viscosity is different under different baking temperatures.⁴⁰ Generally, baked SU-8 has 3 status when temperature increases, solid, glass, and liquid. The corresponding viscosity will decrease and the SU-8 can also have higher fluidity. When the baking temperature is larger than 120 °C, the flow-in speed will increase sharply. But, if the baking temperature is higher, the SU-8 will reflow in the gaps too fast, which makes the flow-in depth hard to be controlled. There is a high chance that the whole gaps will be blocked, and no drugs can flow through these gaps any more. Considering that the total SU-8 supporting structure is only 350 μm high, we choose 125 °C as baking temperature for proper SU-8 flow-in speed and easier SU-8 flow-in depth control.Open in a separate windowFigure 5The relationship between flow-in speed and baking temperature.To ensure the adequate stiffness of the SU-8 microneedles in vertical direction, Instron Microtester 5848 (Instron, USA) was deployed to press the microneedles with the similar method reported by Khoo et al.⁴¹ As shown in Fig. ​Fig.6a,6a, the vertical buckling force was as much as 8.1N, which was much larger than the reported minimal required penetration force.⁴² However, in the previous practical testing experiments, even though the microneedles were strong enough in vertical direction, the planar shear force induced by skin deformation might also break the interface between SU-8 pillars and top tips. In our new device with four pillars supporting structure, the SU-8 could flow inside the sidewall gaps between the pillars to form anchors. These anchors could enhance microneedles'' mechanical strength and overcome the planar shear force problems. Moreover, the anchors strength could be improved by controlling the SU-8 flow-in depth. Fig. ​Fig.77 shows that the flow-in depth increases when the baking time increases as the baking time increases at 125 °C. Fig. ​Fig.6b6b shows that the corresponding planar buckling force can be improved to be larger than 1 N by increasing flow-in depth. Some sidewall gaps at bottom are kept on purpose for drugs delivery; hence, the flow-in depth is chosen as 200 μm.Open in a separate windowFigure 6(a) Measurement of the vertical buckling force. (b) The planar buckling force varies under different flow-in depth (I, II, III, and IV corresponding to the certain images in Fig. ​Fig.77).Open in a separate windowFigure 7Different flow-in depth inside the gaps between SU-8 pillars. (a) 0 μm; (b) 100 μm; (c) 200 μm; and (d) 350 μm (scale bar is 100 μm).The penetration capability of the 3 × 3 SU-8 microneedles array is characterized by conducting the insertion experiment on the porcine cadaver skin. 10 microneedles devices were tested and all of them were strong enough to be inserted into the tissue without any breakage. Histology images of the skin at the site of one microneedle penetration were derived to prove that the sharp conical tip was not broken during the insertion process (Fig. ​(Fig.8).8). It also shows penetrated evidence because the hole shape is the same as the sharp conical tip.Open in a separate windowFigure 8Histology image of individual microneedle penetration (scale bar is 100 μm).In order to verify that the drug solution can be delivered into tissue from the sidewall gaps of the microneedles, FITC (Fluorescein isothiocyanate) (Sigma Aldrich, Singapore) solution was delivered through the SU-8 microneedles after they were penetrated into the mouse cadaver skin. The representative results were then investigated via a confocal microscope (Fig. ​(Fig.9).9). The permeation pattern of the solution along the microchannel created by microneedles confirmed the solution delivery results. The black area was a control area without any diffused florescent solution. In contrast, the illuminated area in Fig. ​Fig.99 indicates the area where the solution has diffused to it. These images were taken consecutively from the skin surface down to 180 μm with 30 μm intervals. The diffusion area had a similar dimension with the inserted microneedles. It has proved that the device can be used to deliver drugs into the body.Open in a separate windowFigure 9Images of confocal microscopy to show the florescent solution is successfully delivered into the tissue underneath the skin surface. (a) 30 μm; (b) 60 μm; (c) 90 μm; (d) 120 μm; (e) 150 μm; and (f) 180 μm (scale bar is 100 μm).Due to the uneven surface of deformed skin, there is always tiny gap happened between tips of some microneedles and local surface skin. The microneedles could not be entirely inserted into the tissue. Drugs might leak to the skin surface through the sidewall gaps under certain driven pressure. Hydrogel absorption experiment was conducted to quantify the delivery rate (i.e., the ratio of solution delivered into tissues in the total delivered volume) and to optimize the delivery speed. Using hydrogel as the tissue model for quantitative analysis of microneedle releasing process was reported by Tsioris et al.⁴³ The details are shown here. Gelatin hydrogel was prepared by boiling 70 ml DI (Deionized) water and mixing it with 7 g of Knox^TM original unflavored gelatin powder. The solution was poured into petri dish to 1 cm high. Then, the petri dish was put into a fridge for half an hour. Gelatin solution became collagen slabs. The collagen slabs were cut into 6 mm × 6 mm sections. A piece of fully stretched parafilm (Parafilm M, USA) was tightly mounted on the surface of the collagen slabs. This parafilm was used here to block the leaked solution further diffusing into the collagen slab in the delivery process. Then, the microneedles penetrated the parafilm and went into the collagen slab. Controlled by a syringe pump, 0.1 ml–0.5 mg/ml glucose solution was delivered into the collagen slab under different speeds. Methylene Blue (Sigma Aldrich, Singapore) was mixed into the solution for better inspection purpose (Fig. 10a). Then, the collagen slabs was digested in 1 mg/ml collagenase (Sigma Aldrich, Singapore) at room temperature (Fig. 10b). It took around 1 h that all the collagen slabs could be fully digested (Fig. 10d). The solution was collected to measure the glucose concentration with glucose detection kit (Abcam, Singapore). Briefly, both diluted glucose standard solution and the collected glucose solution were added into a series of wells in a well plate. Glucose assay buffer, glucose enzyme, and glucose substrate were mixed with these samples in the wells. After incubation for 30 min, their absorbance were examined by using a microplate reader at a wavelength of 450 nm. By comparing the readings with the measured concentration standard curve (Fig. 11a), the glucose concentration in the hydrogel, the glucose absorption rate in the hydrogel, and the solution delivery rate by microneedles could be measured and calculated. As shown in Fig. 11b, when the delivering speed of a single microneedle increased from 0.1 μl/min to 2 μl/min, the glucose absorption rate also increased. Most of the glucose solution from microneedles could go into the hydrogel. The delivered rate could be as high as 71%. The rest solution leaked from sidewall gaps and blocked by parafilm. However, when the delivered speed for a single microneedle was larger than 2 μl/min, the hydrogel absorption rate was saturated. More and more solution could not go into the hydrogel but leak from the sidewall gaps. Then, the delivered rate decreased. Therefore, 2 μl/min was chosen as the optimized delivery speed for the microneedle.Open in a separate windowFigure 10Glucose solution could be delivered into the hydrogel, and the collagen stabs were dissolved by collagenase.Open in a separate windowFigure 11(a) Standard curve for glucose detection; (b) glucose absorption rate and solution delivery rate in a single needle corresponding to different delivery speed.In conclusion, a drug delivery device of integrated vertical SU-8 microneedles array is fabricated based on a new double drawing lithography technology in this study. Compared with the previous biocompatible polymer-based microneedles fabrication technology, the proposed fabrication process is scalable, reproducible, and inexpensive. The fabricated microneedles are rather strong along both vertical and planar directions. It is proved that the microneedles were penetrated into the pig skin easily. The feasibility of drug delivery using SU-8 microneedles is confirmed by FITC fluorescent delivery experiment. In the hydrogel absorption experiment, by controlling the delivery speed under 2 μl/min per microneedle, the delivery rate provided the microneedle is as high as 71%. In the next step, the microneedles will be further integrated with microfluidics on a flexible substrate, forming a skin-patch like drug delivery device, which may potentially demonstrate a self-administration function. When patients need an injection treatment at home, they can easily use such a device just like using an adhesive bandage strip.     

A rapid,inexpensive surface treatment for enhanced functionality of polydimethylsiloxane microfluidic channels

A rapid, inexpensive method using alkoxysilanes has been developed to selectively coat the interior of polydimethylsiloxane (PDMS) microfluidic channels with an integral silicaceous layer. This method combines the rapid prototyping capabilities of PDMS with the desirable wetting and electroosmotic properties of glass. The procedure can be carried out on the open faces of PDMS blocks prior to enclosure of the channels, or by flowing the reagents through the preformed channels. Therefore, this methodology allows for high-throughput processing of entire microfluidic devices or selective modification of specific areas of a device. Modification of PDMS with tetraethoxysilane generated a stable surface layer, with enhanced wettability and a more stable electroosmotic flow rate than native PDMS. Modification of PDMS with 3-aminopropyltriethoxysilane generated a surface layer bearing amine functionalities allowing for further chemical derivatization of the PDMS surface.     

Simulation of malaria-infected red blood cells in microfluidic channels: Passage and blockage

Malaria-infected red blood cells (iRBCs) become less deformable with the progression of infection and tend to occlude microcapillaries. This process has been investigated in vitro using microfluidic channels. The objective of this paper is to provide a quantitative basis for interpreting the experimental observations of iRBC occlusion of microfluidic channels. Using a particle-based model for the iRBC, we simulate the traverse of iRBCs through a converging microfluidic channel and explore the progressive loss of cell deformability due to three factors: the stiffening of the membrane, the reduction of the cell''s surface-volume ratio, and the growing solid parasites inside the cell. When examined individually, each factor tends to hinder the passage of the iRBC and lengthen the transit time. Moreover, at sufficient magnitude, each may lead to obstruction of narrow microfluidic channels. We then integrate the three factors into a series of simulations that mimic the development of malaria infection through the ring, trophozoite, and schizont stages. These simulations successfully reproduce the experimental observation that with progression of infection, the iRBC transitions from passage to blockage in larger and larger channels. The numerical results suggest a scheme for quantifying iRBC rigidification through microfluidic measurements of the critical pressure required for passage.     

On demand nanoliter-scale microfluidic droplet generation,injection, and mixing using a passive microfluidic device

We here present and characterize a programmable nanoliter scale droplet-on-demand device that can be used separately or readily integrated into low cost single layer rapid prototyping microfluidic systems for a wide range of user applications. The passive microfluidic device allows external (off-the-shelf) electronically controlled pinch valves to program the delivery of nanoliter scale aqueous droplets from up to 9 different inputs to a central outlet channel. The inputs can be either continuous aqueous fluid streams or microliter scale aqueous plugs embedded in a carrier fluid, in which case the number of effective input solutions that can be employed in an experiment is no longer strongly constrained (100 s–1000 s). Both nanoliter droplet sequencing output and nanoliter-scale droplet mixing are reported with this device. Optimization of the geometry and pressure relationships in the device was achieved in several hardware iterations with the support of open source microfluidic simulation software and equivalent circuit models. The requisite modular control of pressure relationships within the device is accomplished using hydrodynamic barriers and matched resistance channels with three different channel heights, custom parallel reversible microfluidic I/O connections, low dead-volume pinch valves, and a simply adjustable array of external screw valves. Programmable sequences of droplet mixes or chains of droplets can be achieved with the device at low Hz frequencies, limited by device elasticity, and could be further enhanced by valve integration. The chip has already found use in the characterization of droplet bunching during export and the synthesis of a DNA library.     

Polyelectrolyte multilayer surface functionalization of poly(dimethylsiloxane) (PDMS) for reduction of yeast cell adhesion in microfluidic devices

Polyelectrolyte multilayers (PEMs) based on the combinations poly(diallyldimethylammonium chloride)∕poly(acrylic acid) (PDADMAC∕PAA) and poly(allylamine hydrochloride)∕PAA (PAH∕PAA) were adsorbed on poly(dimethylsiloxane) (PDMS) and tested for nonspecific surface attachment of hydrophobic yeast cells using a parallel plate flow chamber. A custom-made graft copolymer containing poly(ethylene glycol) (PEG) side chains (PAA-g-PEG) was additionally adsorbed on the PEMs as a terminal layer. A suitable PEM modification effectively decreased the adhesion strength of Saccharomyces cerevisiae DSM 2155 to the channel walls. However, a further decrease in initial cell attachment and adhesion strength was observed after adsorption of PAA-g-PEG copolymer onto PEMs from aqueous solution. The results demonstrate that a facile layer-by-layer surface functionalization from aqueous solutions can be successfully applied to reduce cell adhesion strength of S. cerevisiae by at least two orders of magnitude compared to bare PDMS. Therefore, this method is potentially suitable to promote planktonic growth inside capped PDMS-based microfluidic devices if the PEM deposition is completed by a dynamic flow-through process.     

Contactless microfluidic pumping using microchannel-integrated carbon black composite membranes

The ability to pump and manipulate fluid at the micron-scale is a basic requirement for microfluidic platforms. Many current manipulation methods, however, require expensive and bulky external supporting equipment, which are not typically compatible for portable applications. We have developed a contactless metal electro-osmotic micropump capable of pumping conductive buffers. The pump operates using two pairs of gallium metal electrodes, which are activated using an external voltage source and separated from a main flow channel by a thin micron-scale polydimethylsiloxane (PDMS) membrane. The thin contactless membrane allows for field penetration and electro-osmotic flow within the microchannel, but eliminates electrode damage and sample contamination commonly associated with traditional DC electro-osmotic pumps that utilize electrodes in direct contact with the working fluid. Our previous work has demonstrated the effectiveness of this method in pumping deionized water. However, due to the high resistivity of PDMS, this method proved difficult to apply towards manipulating conductive buffers. To overcome this limitation, we fabricated conductive carbon black (CB) powder directly into the contactless PDMS membranes. The increased electrical conductivity of the contactless PDMS membrane significantly increased micropump performance. Using a microfluidic T-channel device and an electro-osmotic flow model, we determined the influence that CB has on pump pressure for CB weight percents varying between 0 and 20. The results demonstrate that the CB increases pump pressure by two orders of magnitude and enables effective operations with conductive buffers.     

High-throughput sorting of drops in microfluidic chips using electric capacitance

We analyze a recently introduced approach for the sorting of aqueous drops with biological content immersed in oil, using a microfluidic chip that combines the functionality of electrowetting with the high throughput of two-phase flow microfluidics. In this electrostatic sorter, three co-planar electrodes covered by a thin dielectric layer are placed directly below the fluidic channel. Switching the potential of the central electrode creates an electrical guide that leads the drop to the desired outlet. The generated force, which deflects the drop, can be tuned via the voltage. The working principle is based on a contrast in conductivity between the drop and the continuous phase, which ensures successful operation even for drops of highly conductive biological media like phosphate buffered saline. Moreover, since the electric field does not penetrate the drop, its content is protected from electrical currents and Joule heating. A simple capacitive model allows quantitative prediction of the electrostatic forces exerted on drops. The maximum achievable sorting rate is determined by a competition between electrostatic and hydrodynamic forces. Sorting speeds up to 1200 per second are demonstrated for conductive drops of 160 pl in low viscosity oil.     

On-chip recalcification of citrated whole blood using a microfluidic herringbone mixer

In vitro assays of platelet function and coagulation are typically performed in the presence of an anticoagulant. The divalent cation chelator sodium citrate is among the most common because its effect on coagulation is reversible upon reintroduction of divalent cations. Adding divalent cations into citrated blood by batch mixing leads to platelet activation and initiation of coagulation after several minutes, thus limiting the time blood can be used before spontaneously clotting. In this work, we describe a herringbone microfluidic mixer to continuously introduce divalent cations into citrated blood. The mixing ratio, defined as the ratio of the volumetric flow rates of citrated blood and recalcification buffer, can be adjusted by changing the relative inlet pressures of these two solutions. This feature is useful in whole blood assays in order to account for differences in hematocrit, and thus viscosity. The recalcification process in the herringbone mixer does not activate platelets. The advantage of this continuous mixing approach is demonstrated in microfluidic vascular injury model in which platelets and fibrin accumulate on a collagen-tissue factor surface under flow. Continuous recalcification with the herringbone mixer allowed for flow assay times of up to 30 min, more than three times longer than the time achieved by batch recalcification. This continuous mixer allows for measurements of thrombus formation, remodeling, and fibrinolysis in vitro over time scales that are relevant to these physiological processes.