Thin film electro-osmotic pumps for biomicrofluidic applications

Electro-osmotic flow (EOF) pumps are attractive for fluid manipulation in microfluidic channels. Open channel EOF pumps can produce high pressures and flow rates, and are relatively easy to fabricate on-chip or integrate with other microfluidic or electrical components. An EOF pump design that is conducive to on-chip fabrication consists of multiple small channel arms feeding into a larger flow channel. We have fabricated this type of pump design using a thin film deposition process that avoids wafer bonding. We have evaluated pumps fabricated on both silicon and glass substrates. Consistent flow rate versus electric field were obtained. For the range of 40–400 V, flow rates of 0.19–2.30 μL∕min were measured. Theoretical calculations of pump efficiency were made, as well as calculations of the mechanical power generated by various pump shapes, to investigate design parameters that should improve future pumps.     

Electroosmotic flow in vapor deposited silicon dioxide and nitride microchannels

Electroosmotic flow was studied in thin film microchannels with silicon dioxide and silicon nitride sidewalls formed using plasma-enhanced chemical vapor deposition (PECVD). A sacrificial etching process was employed for channel fabrication allowing for cross-sections with heights of 3 μm, ranging from 2 μm to 50 μm in width. Flow rates were measured for single channels and multichannel electroosmotic pump structures for pH levels ranging from 2.6 to 8.3, and zeta potentials were calculated for both silicon dioxide and silicon nitride surfaces. Flow rates as high as 0.086 μL∕min were measured for nitride multichannel pumps at applied electric fields of 300 V∕mm. The surface characteristics of PECVD nitride were analyzed and compared to more well-known oxide surfaces to determine the density of amine sites compared to silanol sites.     

Negative dielectrophoretic capture of bacterial spores in food matrices

A microfluidic device with planar square electrodes is developed for capturing particles from high conductivity media using negative dielectrophoresis (n-DEP). Specifically, Bacillus subtilis and Clostridium sporogenes spores, and polystyrene particles are tested in NaCl solution (0.05 and 0.225 S∕m), apple juice (0.225 S∕m), and milk (0.525 S∕m). Depending on the conductivity of the medium, the Joule heating produces electrothermal flow (ETF), which continuously circulates and transports the particles to the DEP capture sites. Combination of the ETF and n-DEP results in different particle capture efficiencies as a function of the conductivity. Utilizing 20 μm height DEP chambers, “almost complete” and rapid particle capture from lower conductivity (0.05 S∕m) medium is observed. Using DEP chambers above 150 μm in height, the onset of a global fluid motion for high conductivity media is observed. This motion enhances particle capture on the electrodes at the center of the DEP chamber. The n-DEP electrodes are designed to have well defined electric field minima, enabling sample concentration at 1000 distinct locations within the chip. The electrode design also facilitates integration of immunoassay and other surface sensors onto the particle capture sites for rapid detection of target micro-organisms in the future.     

An integrated dielectrophoretic chip for continuous bioparticle filtering, focusing, sorting, trapping, and detecting

Multi-target pathogen detection using heterogeneous medical samples require continuous filtering, sorting, and trapping of debris, bioparticles, and immunocolloids within a diagnostic chip. We present an integrated AC dielectrophoretic (DEP) microfluidic platform based on planar electrodes that form three-dimensional (3D) DEP gates. This platform can continuously perform these tasks with a throughput of 3 μL∕min. Mixtures of latex particles, Escherichia coli Nissle, Lactobacillus, and Candida albicans are sorted and concentrated by these 3D DEP gates. Surface enhanced Raman scattering is used as an on-chip detection method on the concentrated bacteria. A processing rate of 500 bacteria was estimated when 100 μl of a heterogeneous colony of 10⁷ colony forming units ∕ml was processed in a single pass within 30 min.     

Complementary metal oxide semiconductor compatible fabrication and characterization of parylene-C covered nanofluidic channels with integrated nanoelectrodes

Nanochannels offer a way to align and analyze long biopolymer molecules such as DNA with high precision at potentially single basepair resolution, especially if a means to detect biomolecules in nanochannels electronically can be developed. Integration of nanochannels with electronics will require the development of nanochannel fabrication procedures that will not damage sensitive electronics previously constructed on the device. We present here a near-room-temperature fabrication technology involving parylene-C conformal deposition that is compatible with complementary metal oxide semiconductor electronic devices and present an analysis of the initial impedance measurements of conformally parylene-C coated nanochannels with integrated gold nanoelectrodes.No two cells are exactly alike, either in terms of their genome, the genomic epigenetic modification of the genome, or the expressed proteome.¹ The genomic heterogeneity of cells is particularly important from an evolutionary perspective since it represents the stages of evolution of a population of cells under stress.² Because of the important variances in the genome that occur from cell to cell, it is critical to develop genomic analysis technologies which can do single-cell and single molecule genomic analysis as an electronic “direct read” without intervening amplification steps.^{3, 4, 5, 6, 7, 8} In this paper, we present a technique which uses conformal coverage of nanochannels containing nanoelectrodes using a room-temperature deposition of parylene-C, a pin-hole-free, excellent electrical insulator with low autofluorescence.⁹ This procedure should open the door to integration of many kinds of surface electronics with nanochannels. One of the most difficult aspects in introducing electronics into nanochannel technology is the sealing of nanochannel so that the electrodes are not compromised by harsh chemicals or high temperatures. There are various methods to form nanochannels containing nanoelectrodes, including wafer bonding techniques,¹⁰ removal of sacrificial materials,¹¹ and nonuniform sputtering deposition.¹² Methods employing a sacrificial layer removal show the greater compatibility to electronic integration, but current methods to remove sacrificial materials require either high temperatures¹¹ or harsh chemicals.^{13, 14}The basic device consisted of 12 mm long, 100 nm wide, 100 nm high nanochannels interrogated by 22 pairs of 30 nm wide gold nanoelectrodes. The outline of the fabrication process is shown in Fig. ​Fig.1.1. The fabrication process was carried out on a standard 4 in. single-side polished p-type ⟨100⟩ silicon wafer with 100 nm of dry thermal oxide on the top as an insulating layer, which also helped the wetting of the nanochannels. The first step involved nanofabrication of the 25 nm thick nanoelectrodes on the SiO₂ top of the wafer using electron beam lithography (EBL). External gold connection pads were constructed using standard metal lift-off techniques and photolithography to connect to the nanoelectrodes. A Raith E-Line e-beam writing system (Raith USA, Ronkonkoma, NY) was used to expose polymethyl methacrylate (PMMA) for metal lift-off. Figure ​Figure1a1a shows a scanning electron microscopy (SEM) image of the nanoelectrodes. The 100 nm sealed nanochannels were constructed using sacrificial removal techniques. We used EBL to expose a 100 nm thick film of PMMA over the gold nanolines in the region around the nanolines, leaving behind lines of unexposed sacrificial layer of PMMA. We next evaporated 25 nm of SiO₂ over the nanolines to improve the surface wetting properties of nanochannel and then conformally coated with 4 μm thick of parylene-C [poly(chloro-p-xylylene)] using a Specialty Coating Systems model PDS 2010 parylene coating system (SCS Systems, Indianapolis, IN). Access holes for the gold electrodes and the feeding channels were etched through by oxygen plasma and 1:10 buffered oxide etchant. To avoid autofluorescence induced in parylene by an active plasma¹⁵ and ambient UV radiation,¹⁶ it is important not to expose the remaining parylene with plasma and to keep the samples in the dark. The sacrificial removal of PMMA in the nanochannels was done in four steps: (1) soaking the chip in 55 °C 1165 MicroChem resist remover (MicroChem, Newton, MA) for 36 h, (2) room-temperature soaking in 1,2-dichloroethane for 12 h, (3) soaking in room-temperature acetone for 12 h, and (4) drying the nanochannels by critical point drying (CPD-030, BAL-TEC AG, Principality of Liechtenstein), which served to prevent the collapse of the nanochannel resulting from surface tension of the acetone.Open in a separate windowFigure 1(a) SEM image of gold nanoelectrodes; scale bar is 200 nm. (b) 100×100 nm² PMMA nanoline is written over the gold nanoelectrodes by exposure of the surrounding PMMA. (c) Parylene-C conformal coating over the PMMA nanoline. PMMA is dissolved and parylene-C etched by reactive ion etching.Conductance measurements were done using ac techniques. The ac impedance Z_tot of an insulating ionic fluid such as water between electrodes is a complex subject.¹⁷ The most general model for the complex impedance of an electrode in ionic solution is typically modeled as the Randle circuit, which is shown in Fig. ​Fig.22.¹⁷ There are two major contributions to the imaginary part of the impedance: the capacitance of the double layer (C_dl), which is purely imaginary and has no dc conductance, and the impedance due to charge transfer resulting in electrochemical reactions at the electrode∕electrolyte interface, which can be modeled as a contact resistor (R_CT), which is given by the Butler–Volmer equation, which describes the I-V characteristic curve when electrochemical reactions occur at the electrode,¹⁸ in series with a complex Warburg impedance (Z_W) which represents injected charge transport near the electrode;¹⁹ more details can be found in Ref. ²⁰. Since we applied a 10 mV rms ac voltage with no dc offset in our measurements, electrochemical reactions are negligible, which means no electrochemical charge transfer occurred and as a result R_CT goes to infinity. We have drawn a gray box around the elements connected to the Warburg impedance branch of the circuit to show that they are negligible in our analysis.Open in a separate windowFigure 2The equivalent circuit of the nanoelectrodes in contact with water lying atop an insulating SiO₂ film which covers a silicon substrate. The elements in the gray boxes can be ignored in our measurements since there is no hydrolysis at low voltage, while the elements within the dotted box are coupling reactances to the underlying p-doped silicon wafer.In the case of no direct charge injection, the electrodes are coupled by the purely capacitive dielectric layer impedance C_dl to the solvent and are also coupled capacitively by the dielectric SiO₂ film capacitance C_ox to the underlying p-doped silicon semiconductor. We model the semiconductor as a purely resistive material with bulk resistivity ρ_Si. The value of C_dl∕area is on the order of ϵϵ_oκ, where ϵ is the dielectric constant of water (about 80) and κ is the Debye screening parameter of the counterions in solution: κ=ϵϵokBT∕e2Σici∞zi2,²⁰ where ci∞ is the bulk ion concentration of charge z_i. At our salt molarity of 50 mM (1∕2 Tris∕Borate∕EDTA (TBE) buffer), C_dl is approximately 30 μF∕cm² using 1∕κ∼1 nm.In Fig. ​Fig.3,3, we show the ac impedance measurements between pairs nanoelectrodes for both dry and TBE buffer wet nanochannels. The electrodes are capacitively coupled to the underlying silicon substrate through an oxide capacitor C_ox. We model the doped silicon wafer as pure resistors, so there is an R₁ that connects both C_ox, and each C_ox is connected to the ground with an R₂. Curve fitting was done by using the 3SPICE circuit emulation code (VAMP Inc., Los Angeles, CA). We therefore obtained the following parameters for the dry curve: C_ox=1.32 nF, R₁=17.5 μΩ, and R₂=32.8 kΩ. R₁ is not sensitive in the fit as long as it is smaller than the impedance of C_ox. Given ρ_Si of the wafer of 1–10 Ω cm, R₂ should be on the order of 10³ Ω, which is slightly smaller than our fitting results. The same parameters for the wafer coupling parameters were then used for fitting the impedance measurements for wet channels. For TBE buffer solution in the nanochannel, curve fitting yields C_dl=50 pF and R_sol=10⁵ Ω. However, given the dimension of our nanochannels, we should get a transverse resistance R∼10⁹ Ω. One possible explanation for this difference is that the evaporated SiO₂ film which was put over the PMMA is porous and allows buffer to penetrate the oxide film,²¹ but given that the film is only 25 nm thick this would at most increase the cross section by one order of magnitude. However, it is known that there is a high fractional presence of mobile counterions associated with the charged channel walls.²² To calculate exact conductance contribution from the surface charges is a tricky business, but since the surface-to-volume ratios in our nanochannels are much greater than the slits, a larger conductance enhancement can be expected, and more work needs to be done.Open in a separate windowFigure 3ac impedance spectra of TBE buffer solution in a transchannel measurement between adjacent pairs of nanoelectrodes separated by 135 μm. The red circles are data for a dry channel and the solid red line is the fit to the model shown in the upper right hand corner. The green squares and dashed green line are for a nanochannel wet with TBE buffer.We have presented a way to fabricate a nanochannel integrated with electrodes. This technology opens up opportunities for electronic detection of charged polymers. With our techniques to fabricate nanoelectrodes with nanochannels, it should be possible to include integrated electronics with nanofludics, allowing the electronic observation of a single DNA molecule at high spatial resolution. However, the present design has problems. Most of the ac went through the silicon wafer instead of the solution. To enhance the sensitivity, we need either to increase the ratio of current going through the liquid to the current going through the wafer or to have a circuit design that picks up the changes in C_dl and R_sol.     

Numerical study of in situ preconcentration for rapid and sensitive nanoparticle detection

This paper presents a numerical study of a preconcentrator design that can effectively increase the binding rate at the sensor in a real time manner. The particle enrichment is realized by the ac electrothermal (ACET) effect, which induces fluid movement to carry nanoparticles toward the sensor. The ACET is the only electrical method to manipulate a biological sample of medium to high ionic strength (>0.1 S∕m, e.g., 0.06× phosphate buffered saline). The preconcentrator consists of a pair of electrodes striding over the sensor, simple to implement as it is electrically controlled. This preconcentrator design is compatible and can be readily integrated with many types of micro- to nanosensors. By applying an ac signal over the electrodes, local vortices will generate a large velocity perpendicular to the reaction surface, which enhances transport of analytes toward the sensor. Our simulation shows that the binding rate at the sensor surface is greatly enhanced. Our study also shows that the collection of analytes will be affected by various parameters such as channel height, inlet velocity, and sensor size, and our results will provide guidance in optimization of the preconcentrator design.     

Fast three dimensional ac electro-osmotic pumps with nonphotolithographic electrode patterning

Three dimensional (3D) stepped electrodes dramatically improve the flow rate and frequency range of ac electro-osmotic pumps, compared to planar electrodes. However, the fabrication of 3D stepped electrodes for ac electro-osmosis (ACEO) pumps usually involves several processing steps. This paper demonstrates results from ACEO pumps produced by a faster and less expensive method to fabricate the 3D electrodes—extending the previous work to disposable devices. The method is based on shadowed evaporation of metal on an insulating substrate that can be injection molded. Flow velocities through the 3D ACEO pump are similar to those seen in the previous work.     

Using nonlinear ac electrokinetics vortex flow to enhance catalytic activities of sol-gel encapsulated trypsin in microfluidic devices

A novel microstirring strategy is applied to accelerate the digestion rate of the substrate N_α-benzoyl-L-arginine-4-nitroanilide (L-BAPA) catalyzed by sol-gel encapsulated trypsin. We use an ac nonlinear electrokinetic vortex flow to stir the solution in a microfluidic reaction chamber to reduce the diffusion length between the immobilized enzyme and substrate in the solution. High-intensity nonlinear electroosmotic microvortices, with angular speeds in excess of 1 cm∕s, are generated around a small (∼1.2 mm) conductive ion exchange granule when ac electric fields (133 V∕cm) are applied across a miniature chamber smaller than 10 μl. Coupling between these microvortices and the on-and-off electrophoretic motion of the granule in low frequency (0.1 Hz) ac fields produces chaotic stream lines to stir substrate molecules sufficiently. We demonstrate that, within a 5-min digestion period, the catalytic reaction rate of immobilized trypsin increases almost 30-fold with adequate reproducibility (15%) due to sufficient stirring action through the introduction of the nonlinear electrokinetic vortices. In contrast, low-frequency ac electroosmotic flow without the granule, provides limited stirring action and increases the reaction rate approximately ninefold with barely acceptable reproducibility (30%). Dye molecules are used to characterize the increases in solute diffusivity in the reaction reservoir in which sol-gel particles are placed, with and without the presence of granule, and compared with the static case. The solute diffusivity enhancement data show respective increases of ∼30 and ∼8 times, with and without the presence of granule. These numbers are consistent with the ratios of the enhanced reaction rate.     

Electrokinetic focusing and filtration of cells in a serpentine microchannel

Focusing cells into a single stream is usually a necessary step prior to counting and separating them in microfluidic devices such as flow cytometers and cell sorters. This work presents a sheathless electrokinetic focusing of yeast cells in a planar serpentine microchannel using dc-biased ac electric fields. The concurrent pumping and focusing of yeast cells arise from the dc electrokinetic transport and the turn-induced ac∕dc dielectrophoretic motion, respectively. The effects of electric field (including ac to dc field ratio and ac field frequency) and concentration (including buffer concentration and cell concentration) on the cell focusing performance were studied experimentally and numerically. A continuous electrokinetic filtration of E. coli cells from yeast cells was also demonstrated via their differential electrokinetic focusing in a serpentine microchannel.     

Optofluidic refractometer using resonant optical tunneling effect

This paper presents the design and analysis of a liquid refractive index sensor that utilizes a unique physical mechanism of resonant optical tunneling effect (ROTE). The sensor consists of two hemicylindrical prisms, two air gaps, and a microfluidic channel. All parts can be microfabricated using an optical resin NOA81. Theoretical study shows that this ROTE sensor has extremely sharp transmission peak and achieves a sensitivity of 760 nm∕refractive index unit (RIU) and a detectivity of 85 000 RIU⁻¹. Although the sensitivity is smaller than that of a typical surface plasmon resonance (SPR) sensor (3200 nm∕RIU) and is comparable to a 95% reflectivity Fabry–Pérot (FP) etalon (440 nm∕RIU), the detectivity is 17 000 times larger than that of the SPR sensor and 85 times larger than that of the FP etalon. Such ROTE sensor could potentially achieve an ultrahigh sensitivity of 10⁻⁹ RIU, two orders higher than the best results of current methods.     

Band-broadening suppressed effect in long turned geometry channel and high-sensitive analysis of DNA sample by using floating electrokinetic supercharging on a microchip

A featured microchip owning three big reservoirs and long turned geometry channel was designed to improve the detection limit of DNA fragments by using floating electrokinetic supercharging (FEKS) method. The novel design matches the FEKS preconcentration needs of a large sample volume introduction with electrokinetic injection (EKI), as well as long duration of isotachophoresis (ITP) process to enrich low concentration sample. In the curved channel [∼45.6 mm long between port 1 (P1) and the intersection point of two channels], EKI and ITP were performed while the side port 3 (P3) was electrically floated. The turn-induced band broadening with or without ITP process was investigated by a computer simulation (using CFD-ACE+ software) when the analytes traveling through the U-shaped geometry. It was found that the channel curvature determined the extent of band broadening, however, which could be effectively eliminated by the way of ITP. After the ITP-stacked zones passed the intersection point from P1, they were rapidly destacked for separation and detection from ITP to zone electrophoresis by using leading ions from P3. The FEKS carried on the novel chip successfully contributed to higher sensitivities of DNA fragments in comparison with our previous results realized on either a single channel or a cross microchip. The analysis of low concentration 50 bp DNA step ladders (0.23 μg∕ml after 1500-fold diluted) was achieved with normal UV detection at 260 nm. The obtained limit of detections (LODs) were on average 100 times better than using conventional pinched injection, down to several ng∕ml for individual DNA fragment.     

Improvement of rectification effects in diffuser/nozzle structures with viscoelastic fluids

This paper reports the improvement of rectification effects in diffuser∕nozzle structures with viscoelastic fluids. Since rectification in a diffuser∕nozzle structure with Newtonian fluids is caused by inertial effects, micropumps based on this concept require a relatively high Reynolds numbers and high pumping frequencies. In applications with relatively low Reynolds numbers, anisotropic behavior can be achieved with viscoelastic effects. In our investigations, a solution of dilute polyethylene oxide was used as the viscoelastic fluid. A microfluidic device was fabricated in silicon using deep reactive ion etching. The microfluidic device consists of access ports for pressure measurement, and a series of ten diffuser∕nozzle structures. Measurements were carried out for diffuser∕nozzle structures with opening angles ranging from 15° to 60°. Flow visualization, pressure drop and diodicity of de-ionized water and the viscoelastic fluid were compared and discussed. The improvement of diodicity promises a simple pumping concept at low Reynolds numbers for lab-on-a-chip applications.     

Unfolding polyelectrolytes in trivalent salt solutions using dc electric fields: A study by Langevin dynamics simulations

We study the behavior of single linear polyelectrolytes condensed by trivalent salt under the action of electric fields through computer simulations. The chain is unfolded when the strength of the electric field is stronger than a critical value. This critical electric field follows a scaling law against chain length, and the exponent of the scaling law is −0.77(1), smaller than the theoretical prediction, −3ν∕2 [R. R. Netz, Phys. Rev. Lett. 90, 128104 (2003)], and the one obtained by simulations in tetravalent salt solutions, −0.453(3) [P.-Y. Hsiao and K.-M. Wu, J. Phys. Chem. B 112, 13177 (2008)]. It demonstrates that the scaling exponent depends sensitively on the salt valence. Hence, it is easier to unfold chains condensed by multivalent salt of a smaller valence. Moreover, the absolute value of chain electrophoretic mobility increases drastically when the chain is unfolded in an electric field. The fact that the mobility depends on electric field and on chain length provides a plausible way to impart chain-length dependence in free-solution electrophoresis via chain unfolding transition induced by electric fields. Finally, we show that, in addition to an elongated structure, a condensed chain can be unfolded into a U-shaped structure. The formation of this structure in our study is purely a result of the electric polarization, not of the elastohydrodynamics dominated in sedimentation of polymers.     

Investigation of microflow reversal by ac electrokinetics in orthogonal electrodes for micropump design

Orthogonal electrodes have been reported to produce high velocity microflows when excited by ac signals, showing potential for micropumping applications. This paper investigates the microflow reversal phenomena in such orthogonal electrode micropumps. Three types of microflow fields were observed by changing the applied electric signals. Three ac electrokinetic processes, capacitive electrode polarization, Faradaic polarization, and the ac electrothermal effect, are proposed to explain the different flow patterns, respectively. The hypotheses were corroborated by impedance analysis, numerical simulations, and velocity measurements. The investigation of microflow reversal can improve the understanding of ac electrokinetics and hence effectively manipulate fluids.     

Ultrafast microfluidics using surface acoustic waves

We demonstrate that surface acoustic waves (SAWs), nanometer amplitude Rayleigh waves driven at megahertz order frequencies propagating on the surface of a piezoelectric substrate, offer a powerful method for driving a host of extremely fast microfluidic actuation and micro∕bioparticle manipulation schemes. We show that sessile drops can be translated rapidly on planar substrates or fluid can be pumped through microchannels at 1–10 cm∕s velocities, which are typically one to two orders quicker than that afforded by current microfluidic technologies. Through symmetry-breaking, azimuthal recirculation can be induced within the drop to drive strong inertial microcentrifugation for micromixing and particle concentration or separation. Similar micromixing strategies can be induced in the same microchannel in which fluid is pumped with the SAW by merely changing the SAW frequency to rapidly switch the uniform through-flow into a chaotic oscillatory flow by exploiting superpositioning of the irradiated sound waves from the sidewalls of the microchannel. If the flow is sufficiently quiescent, the nodes of the transverse standing wave that arises across the microchannel also allow for particle aggregation, and hence, sorting on nodal lines. In addition, the SAW also facilitates other microfluidic capabilities. For example, capillary waves excited at the free surface of a sessile drop by the SAW underneath it can be exploited for micro∕nanoparticle collection and sorting at nodal points or lines at low powers. At higher powers, the large accelerations off the substrate surface as the SAW propagates across drives rapid destabilization of the drop free surface giving rise to inertial liquid jets that persist over 1–2 cm in length or atomization of the entire drop to produce 1–10 μm monodispersed aerosol droplets, which can be exploited for ink-jet printing, mass spectrometry interfacing, or pulmonary drug delivery. The atomization of polymer∕protein solutions can also be used for the rapid synthesis of 150–200 nm polymer∕protein particles or biodegradable polymeric shells in which proteins, peptides, and other therapeutic molecules are encapsulated within for controlled release drug delivery. The atomization of thin films behind a translating drop containing polymer solutions also gives rise to long-range spatial ordering of regular polymer spots whose size and spacing are dependent on the SAW frequency, thus offering a simple and powerful method for polymer patterning without requiring surface treatment or physical∕chemical templating.     

Enhancement of biosensing performance in a droplet-based bioreactor by in situ microstreaming

A droplet-based micro-total-analysis system involving biosensor performance enhancement by integrated surface-acoustic-wave (SAW) microstreaming is shown. The bioreactor consists of an encapsulated droplet with a biosensor on its periphery, with in situ streaming induced by SAW. This paper highlights the characterization by particle image tracking of the speed distribution inside the droplet. The analyte-biosensor interaction is then evaluated by finite element simulation with different streaming conditions. Calculation of the biosensing enhancement shows an optimum in the biosensor response. These results confirm that the evaluation of the Damköhler and Peclet numbers is of primary importance when designing biosensors enhanced by streaming.It has been pointed out that biosensing performances can be limited by the diffusion of the analytes near the sensing surface.¹ In the case of low Peclet number hydrodynamic flows, typical of microfluidic systems, molecule displacements are mainly governed by diffusive effects that affect time scales and sensitivity. To overcome this problem, the enhancement of biosensor performance by electrothermal stirring within microchannels was first reported by Meinhart et al.² Other authors^{3, 4} numerically studied the analyte transport as a function of the position of a nanowire-based sensor inside a microchannel, stressing on the fact that the challenge for nanobiosensors is not the sensor itself but the fluidic system that delivers the sample. Addressing this problem, Squires et al.⁵ developed a simple model applicable to biosensors embedded in microchannels. However, the presented model is limited to the case of a steady flow. The use of surface-acoustic waves (SAWs) for stirring in biomicrofluidic and chemical systems is becoming a popular investigation field,^{6, 7, 8, 9} especially to overcome problems linked to steady flows by enhancing the liquid∕surface interaction.^{1, 10, 11} The main challenges that need to be addressed when using SAW-induced stirring are the complexity of the flow and its poor reproducibility. However, some technical solutions were proposed to yield a simplified microstreaming. Yeo et al. presented a centrifugation system based on SAW that produces the rotation of the liquid in a droplet in a reproducible way by playing on the configuration of the transducers and reflectors,¹² and presented a comprehensive experimental study of the three-dimensional (3D) flow that causes particle concentration in SAW-stirred droplets,¹³ revealing the presence of an azimuthal secondary flow in addition to the main vortexlike circular flow present in acoustically stirred droplets. The efficiency of SAW stirring in microdroplets to favorably cope with mass transport issues was finally shown by Galopin et al.,¹⁴ but the effect of the stirring on the analyte∕biosensor interaction was not studied. It is expected to overcome mass transport limitations by bringing fresh analytes from the bulk solution to the sensing surface.The studied system, described in Fig. ​Fig.1,1, consists of a microliter droplet microchamber squeezed between a hydrophobic piezoelectric substrate and a hydrophobic glass cover. Rayleigh SAWs are generated using interdigitated transducers (interdigital spacing of 50 μm) laid on an X-cut LiNbO₃ substrate.^{1, 15, 16} The hydrophobicity of the substrate and the cover are obtained by grafting octadecyltrichlorosilane (OTS) self-assembled monolayers (contact angle of 108° and hysteresis of 9°). To do so, the surface is first hydroxylized using oxygen plasma (150 W, 100 mT, and 30 sccm³ O₂) during 1 min and then immersed for 3 h into a 1 mM OTS solution with n-hexane as a solvent.Open in a separate windowFigure 1(a) General view of the considered system. (b) Mean value of the measured speeds within the droplet as a function of the inlet power before amplification.When Rayleigh waves are radiated toward one-half of the microchamber, a vortex is created in the liquid around an axis orthogonal to the substrate due to the momentum transfer between the solid and the liquid. This wave is generated under the Rayleigh angle into the liquid.Speed cartographies of the flow induced in the droplet are realized using the particle image tracking technique for different SAW generation powers. To do so, instantaneous images of the flow are taken with a high-speed video camera at 200 frames∕s and an aperture time of 500 μs on a 0.25 μl droplet containing 1 μm diameter fluorescent particles. Figure ​Figure11 shows the mean speed measured in the droplet as a function of the inlet power. The great dependence of the induced mean speed with the SAW power enables a large range of flow speeds in the stirred droplet. Moreover, the flow was visualized with a low depth of field objective. It was found to be circular and two dimensional (2D) in a large thickness range of the droplet.The binding of analytes to immobilized ligands on a biosensor is a two step process, including the mass transport of the analyte to the surface, followed by a complexation step,Abulk↔kmAsurface+B↔ka,kdAB(1)with k_m as the constant rate for mass transport from and to the sensor, and k_a and k_d as the constant rates of association and dissociation of the complex.At the biosensor surface, the reaction kinetics consumes analytes but their transport is limited by diffusive effects. In this case, the Damköhler number brings valuable information by comparing these two effects. Calling the characteristic time of reaction and diffusion, respectively, τ_C and τ_M, the mixing time in diffusion regime can be approximated by τ_M≈h²∕D with D as the diffusion coefficient and h a characteristic length of the microchannel. Calling R_T the ligand concentration on the surface in mole∕m², the Damköhler number (D_a) can be written asDa=τMτC=kaRThD.(2)Depending on the type of reaction, the calculation of D_a helps determine if a specific biointeraction will benefit from a mass SAW-based microstreaming. If the Damköhler number is low, the reaction is slow compared to mass transport and the reaction will not significantly benefit from microstirring. For example, the hybridization of 19 base single stranded DNA in a microfluidic system with a characteristic length of 500 μm is characterized by a Damköhler number of 0.07 and is therefore not significantly influenced by mass transport. On the contrary, the binding of biotin to immobilized streptavidin is characterized by a D_a number of approximately 10⁴. In this case, the stirring solution will significantly improve the reaction rate.COMSOL numerical simulations were carried out to study the efficiency of the SAW stirring in the case of a droplet-based microbioreactor with a diameter of 1 mm. Assuming a 2D flow, the simulated model takes into account the convective and diffusive effects in the analyte-carrying fluid and the binding kinetics on the biosensor surface. This approach was thoroughly developed by Meinhart et al.²On the biosensor surface, the following equations are solved:∂B∂t=kacs(RT−B)−kdB,(3)∂B∂t=D|∂c∂y|y=0(4)with c as the local concentration of analytes in the droplet and B as the surface concentration of bound analytes on the biosensor surface. Simulation results show that a depleted zone is formed near the biosensor in the case of an interaction without stirring. This zone is characterized by a low concentration of analytes and results from the trapping of analytes on the biosensor surface, thus creating a concentration gradient on the vicinity of the biosensor. When stirring is applied, the geometry of the depleted zone is modified, as it is pushed in the direction of the flow. The geometry of the depleted zone then depends on many parameters, among which the diffusion coefficient D, the speed distribution of the flow (not only near the biosensor but also in the whole microfluidic system), and the reaction kinetics on the biosensor. In our case, which is assimilated to a simple circular flow, the depleted zone reaches a permanent state consisting of an analyte-poor layer situated in the exterior perimeter of the stirred droplet. The diffusion of analytes is then limited again by diffusion from the inner part of the droplet toward its exterior perimeter (see Fig. ​Fig.22).Open in a separate windowFigure 2(a) Mean concentration of bound analytes vs time for different mean flow speeds. (b) The obtained concentration profiles with and without circular stirring, t=10 000 s.The initial analyte and receptor concentrations are, respectively, 0.1 nM in the solution and 3.3×10⁻³ nM m on the biosensor surface, the diffusion coefficient is D=10⁻¹¹ m² s⁻¹, and the reaction constants are k_a=10⁶ M⁻¹ s⁻¹ and k_d=10⁻³ s⁻¹. Simulations show that the mean concentration of bound analytes highly increases with the flow speed, improving the efficiency of the biosensing device. To evaluate the benefits of in situ microstreaming with SAW, the same simulations were conducted for D_a numbers ranging from 10⁴ to 10⁸ M⁻¹∕s, by ranging the diffusion coefficient from 4×10⁻¹² to 4×10⁻⁹ m²∕s, and the association coefficient k_a from 10⁴ to 10⁸ M⁻¹∕s. The enhancement factor of analyte capture, defined as the ratio of the binding rate with streaming B and the binding rate without streaming B0, is plotted in Fig. ​Fig.33 for different values of D_a. Calculations are done in the case of a mean flow speed of 0.5 mm∕s.Open in a separate windowFigure 3(a) Enhancement factor (defined as the ratio between binding rate with streaming B and binding rate without streaming B0) for different Damkhöler numbers and (b) normalized enhancement factor for different Peclet numbers.One can notice the saturation of the enhancement factor curve for large value of D_a to the value of 3.5 for high D_a. This can be explained by the fact that for large k_a∕D_a ratios, the analytes, which normally require penetration in the depleted zone by diffusion, do not have time to interact with the biosensor when they pass in the vicinity of its surface. The efficiency of the streaming is then reduced for large values of D_a. In the case of our specific flow configuration, the enhancement factor reaches 3.2 for the interaction of streptavidin on immobilized biotin (D_a=10³).The reported simulation results can be compared to an experimental value obtained using the droplet-based surface plasmon resonance sensor streamed in situ using SAW reported by Yeo et al.¹² By monitoring the streptavidin∕biotin binding interaction on an activated gold slide, they showed that SAW stirring brings an improvement factor of more than 2. This difference can be accounted to the high complexity of the induced 3D flow, which was modeled in a simple manner in our calculations.Other factors must be taken into account when optimizing the improvement factor, such as the flow velocity and the characteristic length of the mixing. To do so, the Peclet number allows the comparison of the convective and diffusive effects.¹⁷ For δC a typical variation in concentration on the distance h, the Peclet number is given byPe=UhD.(5)A significantly high Peclet number causes a decrease in biosensing efficiency as the analytes do not have enough time to interact with the biosensing surface by diffusion through the analyte-poor layer. On the contrary, the case of a low Peclet number corresponds to the diffusion-limited problem. Therefore, for each Damköhler number, there is a Peclet number optimizing this factor. To illustrate this fact, Fig. ​Fig.3b3b shows the calculation of the enhancement factor as a function of the Peclet number for a given D_a.In this paper, we showed that surface loading of typical analytes on a droplet-based biosensor can be highly increased by SAW microstirring. The system permits the enhancement of the biosensing performances by the continuous renewal of the analyte-carrying fluid near the sensing surface. Thanks to mean flow speeds measured up to 1800 μm∕s, the SAW microstreaming can be beneficial to the biosensing of a large range of analyte∕ligand interactions. In addition to the biosensing performance improvement, such a method can be easily integrated in micro-micro-total-analysis systems, which makes it a convenient tool for liquid handling in future biochips.     

AC-dielectrophoretic characterization and separation of submicron and micron particles using sidewall AgPDMS electrodes

The recent development of microfluidic “lab on a chip” devices requires the need to continuously separate submicron particles. Here, we present a PDMS microfluidic device with sidewall conducting PDMS (AgPDMS) composite electrodes capable of separating submicron particles in hydrodynamic flow. In particular, the device can service dual functions. First, the AgPDMS composite electrodes embedded in a sidewall of the device channel allow for performing AC-dielectrophoretic (DEP) characterization through direct microscopic observation of particle behavior. Characterization experiments are carried out for numerous parameters including particle size, medium conductivity, and AC field frequency to reveal important dielectrophoresis DEP information in terms of the crossover frequency and positive/negative DEP behavior under specific frequencies. Second, the device offers an advantage that sidewall AgPDMS composite electrodes can produce strong DEP effects throughout the entire channel height, and thus the robustness of the on-chip particle separation is demonstrated for continuous separation in a flowing mixture of 0.5 and 5 μm particles with 100% separation efficiency.     

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.     

Electrophoretic separation of neurotransmitters on a polystyrene nano-sphere�Mpolystyrene sulphonate coated poly(dimethylsiloxane) microchannel

In this paper, a poly(dimethylsiloxane) microchip with amperometric detector was developed for the electrophoretic separation and determination of neurotransmitters. For increasing the separation efficiency, the microchannel is modified by polystyrene sulphonate∕polystyrene nano-sphere self-assembly coating. A stable electro-osmotic flow (EOF) and higher separation efficiency are obtained in proposed modified microchannel. Under optimized conditions, dopamine, epinephrine, catechol, and serotonin are acceptably baseline separated in this 3.5 cm length separation channel with the theoretical plate number from 4.6 × 10⁴ to 2.1 × 10⁵ per meter and resolution from 1.29 to 12.5. The practicability of proposed microchip is validated by the recovery test with cerebrospinal fluid as real sample which resulted from 91.7% to 106.5%.     

Syringe-vacuum microfluidics: A portable technique to create monodisperse emulsions

We present a simple method for creating monodisperse emulsions with microfluidic devices. Unlike conventional approaches that require bulky pumps, control computers, and expertise with device physics to operate devices, our method requires only the microfluidic device and a hand-operated syringe. The fluids needed for the emulsion are loaded into the device inlets, while the syringe is used to create a vacuum at the device outlet; this sucks the fluids through the channels, generating the drops. By controlling the hydrodynamic resistances of the channels using hydrodynamic resistors and valves, we are able to control the properties of the drops. This provides a simple and highly portable method for creating monodisperse emulsions.Droplet-based microfluidic devices use micron-scale drops as “test tubes” for biological reactions.^{1, 2, 3} With the devices, the drops are loaded with cells, incubated to stimulate cell growth, picoinjected to introduce additional reagents, and sorted to extract rare specimens.^{4, 5, 6} This allows biological reactions to be performed with greatly enhanced speed and efficiency over conventional approaches: by reducing the drop volume, only picoliters of reagent are needed per reaction, while through the use of microfluidics, the reactions can be executed at rates exceeding hundreds of kilohertz. This combination of incredible speed and efficient reagent usage is attractive for a variety of applications in biology, particularly those that require high-throughput processing of reactions, including cell screening, directed evolution, and nucleic acid analysis.^{7, 8} The same advantages of speed and efficiency would also be beneficial for applications in the field, in which the amount of material available for testing is limited, and results are needed with short turnaround. However, a challenge to using these techniques in field applications is that the control systems developed to operate the devices are intended for use in the laboratory: to inject fluids, mechanical pumps are needed, while computers must adjust flow rates to maintain optimal conditions in the device.^{9, 10, 11, 12} In addition to significantly limiting the portability of the system, these qualities make them impractical for use outside the laboratory. For droplet-based microfluidic techniques to be useful for applications in the field, a general, robust, and portable system for operating them is needed.In this paper, we introduce a general, robust, and portable system for operating droplet-based microfluidic devices. In this system, which we call syringe-vacuum microfluidics (SVM), we load the reagents needed for the emulsion into the inlets of a microfluidic drop maker; using a standard plastic syringe, we generate a vacuum at the outlet of the drop maker,¹³ sucking the reagents through the channels, generating drops, and transporting them to different regions for visualization and analysis. By controlling the vacuum strength and channel resistances using hydrodynamic resistors^{14, 15, 16} and single-layer membrane valves,^{17, 18} we are able to specify the flow rates in different regions of the device and to adjust them in real time. No pumps, control computers, or electricity is needed for these operations, making the entire system portable and of potential use for field applications. To characterize the adjustability and precision of this system, we vary channel resistances and vacuum pressures while measuring the effects on drop size and production frequency. We also show how to use this to form drops of many distinct reagents simultaneously using only a single vacuum syringe.Monodisperse drop formation is the central operation in droplet-based microfluidics but can be quite challenging due to the need for precise, steady pumping of reagents; forming monodisperse drops with controlled properties is thus a stringent demonstration of the effectiveness of a control system. While there are many geometries available for microfluidic drop formation,¹⁹ in this discussion we use a simple cross-junction for its proven ability to form uniform emulsions at high rates of speed,^{20, 21} a schematic of which is shown in Fig. ​Fig.1.1. The devices are fabricated in poly(dimethylsiloxane) (PDMS) using soft lithography.²² The drop formation channels have dimensions of 25 μm in width and 25 μm in height. To enable production of aqueous drops in oil, which are the most useful for biological assays, we require hydrophobic devices, which we achieve using an Aquapel chemical treatment: we flow Aqualpel through the channels for a few seconds, flush with air, and then bake the devices for 20 min at 65 °C. After this treatment, the channels are permanently hydrophilic, as is needed for forming aqueous-in-oil emulsions. To introduce reagents into the device, we use 200 μl plastic pipette tips inserted into the channel inlets. To apply the suction, we use a 10 ml Bectin-Dickenson plastic syringe coupled to the device through a 16 G needle and PE∕5 tubing. The other end of the tubing is inserted into the outlet of the device.Open in a separate windowFigure 1Schematic of the microfluidic drop maker for use with SVM. To form water drops in oil, the device must be hydrophobic, which we achieve by treating the channels with Aquapel. The water and surfactant-containing oil are loaded into pipette tips inserted into the device inlets at the locations indicated. To pump the fluids through the drop maker, a syringe applies a vacuum to the outlet; this sucks the fluids through the drop maker, forming drops. The drops are collected into the suction syringe, where they can be stored, incubated, and reintroduced into a microfluidic device for additional processing.To begin forming drops, we fill the device with HFE-7500 fluorocarbon oil, displacing trapped air bubbles that could restrict flow and interfere with drop formation. Pipette tips containing reagents are then inserted into the device inlets, as shown in Fig. ​Fig.11 and pictured in Fig. ​Fig.2a;2a; during this step, care must be taken to not trap air bubbles under the pipette tips, as they would restrict flow. For the fluids, we use distilled water for the droplet phase and HFE-7500 with the ammonium salt of Krytox 157 FSL at 1.8 wt % for the continuous phase. The suction syringe is then connected to the device outlet; to initiate drop formation, the piston is pulled outward and locked in place with a 1 in. binder clip, as shown in Fig. ​Fig.2a.2a. This expands the air in the syringe, generating a vacuum that is transferred to the device through tubing. Since the inlet reagents are open to the atmosphere and thus maintained at a pressure of 1 atm, this creates a pressure differential through the device that pumps the fluids. As the fluids flow through the cross-channel, forces are generated that create drops, as shown in Fig. ​Fig.2b2b (enhanced online). Due to the very steady flow, the drops are highly monodisperse, as shown in Fig. ​Fig.2c.2c. After they are formed, the drops flow out of the device through the suction tube and are collected into the syringe. Depending on the emulsion formulation, drops may coalesce on the metal needle of the syringe; if so, an Upchurch fitting should be used to couple the tubing instead. The collected drops can be stored in the syringe, incubated, and reintroduced into additional microfluidic devices, as needed for the assay.Open in a separate windowFigure 2Photograph of the microfluidic drop formation device with pipette tips containing emulsion reagents and vacuum syringe for pumping (a). Distilled water is used for the droplet phase and HFE-7500 fluorocarbon oil with fluorinated surfactant for the continuous phase. The vacuum applies a pressure differential through the device that pumps the fluids through the drop maker (b) forming drops. The drops are monodisperse, due to the controlled properties of drop formation in microfluidics (c). The scale bars denote 50 μm (enhanced online).In many biological applications, drop size must be precisely controlled. This is essential, for example, when encapsulating molecules or cells in the drops, in which the number encapsulated depends on the drop size.^{3, 23, 24} With SVM, the drop size can be precisely controlled. Our strategy to accomplish this is motivated by the physics of microfluidic drop formation. In microfluidic devices, the capillary number of the flow is normally small, Ca<0.1; as a consequence, the drop formation physics follows a plugging∕squeezing mechanism, in which the drop size depends on the flow rate ratio of the dispersed-to-continuous phase.^{20, 25} By adjusting this ratio, we can thus control the drop size. To adjust this ratio, we use hydrodynamic resistor channels.^{14, 15, 16} These channels are analogous to electronic resistors in that for a fixed pressure drop (voltage) the flow rate through them (current) is inversely proportional to their resistance. By making the resistors longer or shorter, we adjust their resistance, thereby controlling the flow rate.To use resistors to control the drop size, we place three on the inlets of the cross-junction, at the locations indicated in Fig. ​Fig.3a.3a. In this configuration, the flow rate ratio depends on the resistances of the central and side resistors: shortening the side resistors increases the continuous phase flow rate with respect to the dispersed phase, thereby reducing the ratio and, consequently, the drop size, whereas lengthening it increases the drop size. By varying the ratio, we produce drops over a range of sizes, as shown in Fig. ​Fig.3b3b (enhanced online). The drop size is linear in the resistance ratio, indicating that it is linear in the flow rate ratio, as is expected for plugging∕squeezing drop formation [Fig. ​[Fig.3b3b].^{20, 25} This behavior is identical to that of pump-driven fluidics, demonstrating that SVM affords similar control.Open in a separate windowFigure 3Drop properties can be controlled using resistor channels. The resistors are placed on the inlets of the drop maker at the locations indicated in (a). The resistors enable the flow rates of the inner and continuous phases to be controlled. By varying the length ratio of the inlet resistors, we control the flow rate ratio in the drop maker. This allows the drop volume to be controlled, as shown by drop volume plotted as a function of inlet resistor length ratio in (b); varying this ratio does not significantly affect the drop formation frequency, as shown in (c). By varying the length of the outlet resistor, we control the total flow rate through the device; this allows us to form drops of constant volume, but at a different formation frequency, as shown by the plots of volume and frequency as a function of the inverse of the outlet resistor length in (d) and (e), respectively. The measured hydrodynamic resistance of a resistor channel with water as a function of length is shown as inset into (d) (enhanced online).We can also control the frequency of the drop formation using resistor channels. We place a resistor on the outlet of the device; this sets the total flow rate through the device, thereby adjusting drop frequency, as shown in Fig. ​Fig.3e3e (enhanced online). To confirm that the size and frequency control are independent, we plot size as a function of the outlet resistance and frequency as a function of the resistance ratio [Figs. ​[Figs.3c,3c, ​,3d];3d]; both are constant as a function of these parameters, again demonstrating independent control. Frequency can also be adjusted by changing the strength of the vacuum, which can be accomplished by loading a prescribed volume of air into the syringe before expansion. In this case, the vacuum pressure applied is P_fin=V_in∕V_fin×P_in, where V_in is the initial volume of air in the syringe, V_fin is the volume after expansion, and P_in is the initial pressure, which is 1 atm. By loading a prescribed volume of air into the syringe before connecting it to the device and pulling the piston, the expansion factor can be reduced, thereby lowering the vacuum strength.The flow rates through the microfluidic device depend on the applied pressure differential, which, in turn, depends on the value of the ambient pressure. Since ambient pressure may vary due to differences in altitude, the drop formation may also vary. However, since ambient pressure variations affect the inner and outer phase flows equally, this should alter the total flow rate but not the flow rate ratio. Consequently, we expect it to alter drop formation frequency but not drop size because while the frequency depends on absolute flow rate [as illustrated by Fig. ​Fig.3e],3e], drop size depends on the flow rate ratio [as illustrated in Fig. ​Fig.3b].3b]. Based on normal variations in atmospheric pressure on the surface of the Earth, we expect this to produce differences in the drop formation frequency of ∼25%, for example, when operating a device at sea level compared to at the top of a moderately sized mountain.Resistor channels allow drop properties to be controlled, equivalent to what is possible with pump-driven flow; however, they do not allow real-time control because their dimensions are fixed during the fabrication. Real-time control is often needed, for example, as it is when performing reactions in drops for the first time, in which the optimal drop size is not known. To enable real-time control, we must adjust flow rates, which can be achieved using the fluidic analog of electronic potentiometers. Single-layer membrane valves are analogous fluidic components, consisting of a control channel that abuts a flow channel.^{17, 18} By pressurizing the control channel, the thin PDMS membrane between these channels is deflected laterally, constricting the flow channel, thereby increasing its hydrodynamic resistance and reducing its flow rate.¹⁸ To use these membrane valves to vary drop size, we replace the inlet resistors with inlet valves, as shown in Fig. ​Fig.4a.4a. To set the flow rate through a path, we actuate the valve with a defined pressure. To actuate the valves, we use air-filled syringes: a 1 ml syringe is filled with air and connected to the valve control channel through tubing; an additional component, a three-way stopcock is inserted between the syringe and needle, allowing the pressure to be locked in after optimal actuation conditions are obtained. We use one syringe to control the dispersed phase valves and another to control the continuous phase valves. The valves are pressurized by compressing the air in the syringes to a defined degree using the marked graduations; this is achieved by pressing the piston to a defined graduation mark, compressing the air contained within it, thus increasing pressure. The stopcock is then switched to the off position, locking in the actuation. This simple scheme allows precise actuation of the valves, for accurate, defined flow rates in the drop maker, and controlled drop size, as shown in Figs. ​Figs.4b,4b, ​,4c4c (enhanced online). The drop size can be varied at a rate of several hertz without noticeable loss of control; moreover, changing the drop size does not affect the frequency, indicating that, again, these properties are independent, as shown by the constant drop frequency with varying pressure ratio in Fig. ​Fig.4d4d.Open in a separate windowFigure 4Single-layer membrane valves allow the drop size to be varied in real time to screen for optimal reaction conditions. The valves are positioned on the inner and side inlets, as indicated in (a). By adjusting the actuation pressures of the valves, we vary the flow rates in the drop maker, thereby changing the drop size (b), as shown by the plot of drop volume as a function of the actuation pressure ratio in (c). Varying the inlet resistance ratio does not significantly alter drop formation frequency, as shown by frequency as a function of the pressure ratio in (d). A movie of drop formation during actuation of the valves are available in the supplemental material (Ref. ²⁹). The scale bars denote 100 μm (enhanced online).Another useful attribute of SVM is that it readily lends itself to parallel drop formation²⁶ because the pressure that pumps the fluids through the channels is supplied by the atmosphere and is applied evenly over the whole outer surface of the device. This allows fluids to be introduced at equal pressures from different inlets, for forming drops with identical properties in different drop makers. To illustrate this, we use a parallel drop formation device to emulsify eight distinct reagents simultaneously; the product of this is an emulsion library, consisting of drops of identical size in which different drops encapsulate distinct reagents, useful for certain biological applications of droplet-based microfluidics.⁷ The microfluidic device consists of eight T-junction drop makers.²⁵ The drop makers share one oil inlet and outlet but each has its own inner-phase inlet, as shown in Fig. ​Fig.5.5. The oil and outlet channels are wide, ensuring negligible pressure drop through them, so that all T-junctions are operated under the same flow conditions. A distinct reagent fluid is introduced into the inner phase of each T-junction, for which we use eight concentrations of the dye Alexa Fluor 680 in water. After loading these solutions into the device through pipette tips, a syringe applies the vacuum to the outlet, sucking the reagents through the T-junctions, forming drops, as shown by the magnified images of the T-junctions during drop formation in Fig. ​Fig.5.5. Since the drop makers are identical and operated under the same flow conditions, the drops formed are of the same size, as shown in the magnified images in Fig. ​Fig.55 and in a movie available in the supplemental material.²⁹Open in a separate windowFigure 5Parallel drop formation device consisting of eight T-junction drop makers. The drop makers share a common oil inlet and outlet, both of which are wide to ensure even pressure distribution to all drop makers; support posts prevent these channels from collapsing under the suction. Each drop maker has its own inner-phase inlet, allowing emulsification of a distinct reagent. Since the drop maker dimensions and pressure differentials are constant through all drop makers, the drops formed are of the same size, as shown in the magnified images. The drops are ∼35 μm in diameter.To verify that the dye solutions are successfully encapsulated, we image a sample of the collected drops with a fluorescent microscope. The drops are confined in a monolayer between two glass plates so they can be individually imaged. They are of the same size but have distinct fluorescence intensities, as shown in Fig. ​Fig.6a.6a. To quantify these differences, we measure the intensity of each drop and plot the results as a histogram [see Fig. ​Fig.6b].6b]. There are eight peaks in the histogram, corresponding to the eight dye concentrations, demonstrating that all dyes are encapsulated successfully. The peak areas are also similar, demonstrating that drops of different types are formed in equal amounts due to the uniformity of the parallel drop formation.Open in a separate windowFigure 6Fluorescent microscope image of emulsion library created with parallel T-junction device (a). In this demonstration, eight concentrations of Alexa Fluor 680 dye are emulsified simultaneously, producing an emulsion library of eight elements. The drops are of the same size but encapsulate distinct concentrations of the dye solution, as demonstrated by the eight peaks in the intensity histograms in (b). The scale bar denotes 100 μm.SVM is a simple, accessible, and highly controlled way to form monodisperse emulsions for biological assays. It allows controlled amounts of different reagents to be encapsulated in individual drops, drop size to be precisely controlled, and the ability to form drops of different reagents at the same time, in a parallel drop formation device. These properties should make SVM useful for biological applications of monodisperse emulsions;^{1, 2, 3} the portability of SVM should also make it useful for applications in the field, particularly when no electrical power source is available. The parallel emulsification technique should also be useful for particle templating from drops, in which the particles must be of the same size but composed of distinct materials.^{26, 27, 28, 29}