Heusler合金Ni-Mn-Ga和NaZn_(13)型化合物La(Fe,Co,M)_(13),M=Si,Al的巨大磁熵变(英)     

胡凤霞  沈保根 《中国科学院研究生院学报》2002,(2)

在Heusler合金Ni Mn Ga和NaZn1 3型化合物La(Fe,Co,M)1 3,M =Si,Al中很宽温度区间发现巨大磁熵变△S。这两种材料在室温区的磁熵变△S均显著地超过单质稀土Gd并达到著名的磁热效应材料Gd5Si2 Ge2 合金的磁熵变幅度.Heusler合金Ni Mn Ga中的巨大磁熵变来源于具有一级相变特征的马氏结构相变。NaZn1 3型化合物La(Fe,Co,M)1 3,M =Si,Al中异常巨大的磁熵变与合金中的强的磁晶耦合相关,表现为居里温度附近晶格的巨大负膨胀。材料的低价格和其巨大磁熵变表明,Heusler合金Ni Mn Ga和NaZn1 3型化合物La(Fe,Co,M)1 3,M =Si,Al在很宽的高温区,尤其在室温区作为磁制冷工质非常有吸引力。  相似文献   

超微LaMO3（M=Fe,Cr,Mn,Co）的导电性与粒度的关系   总被引：1，自引：0，他引：1  

董相廷  郭奕柱 《科技通报》1994,10(5):277-280

用Ｓｏｌ－Ｇｅｌ法合成了ＬａＭＯ＿３（Ｍ＝Ｆｅ，Ｃｒ，Ｍｎ，Ｃｏ）超微粉末．Ｘ射线衍射分析表明，不同焙烧温度下所合成的化合物均为单相．ＳＥＭ分析表明，粒子呈球形，室温电阻率测量表明，电阻率随粒度的变化呈两种类型：ＬａＦｅＯ＿３，ＬａＣｒＯ＿３随粒度的增大，电阻率增加；而ＬａＭｎＯ＿３，ＬａＣｏＯ＿３随粒度的增大，电阻率减小．  相似文献   

百合科细胞分类学研究——(2)浙江产8属8种的染色体数目和核型报道     

傅承新  洪德元 《中国科学院研究生院学报》1989,27(6):439-450

本文对浙江产的Disporum sessile， Tricyrtis macropodo，  Scilla rcilloides，Ophiopogon japoni-   cus,Liriope platyphylla和Allium macrostemon的核型作了分析，并报道了Polygonatum odoratum   和Asparagus cochinchinensis 的单倍体染色体数目。其中绝大多数为国产材料的首次记载。  相似文献   

浙江产7种菝葜的染色体研究     

傅承新  洪德元 《中国科学院研究生院学报》1990,28(3):211-222

本文报道浙江产菝葜属smilax 7个种的染色体数目和核型。S．nipponica有两种核型，2n＝   26和2n=32，均为3B型，但后一种细胞型的雄株的第一对染色体大小不等，可能为性染色体；S.   riparia 2n=30，属3B型；S．siebodii n＝16；S. china有两个染色体数目，2n＝96 和n＝15；   S. davidiana 2n=32，属3B型，对减数分裂MI的观察发现n＝16；S．glabra 2n=32，亦属3B   型：S. nervo-marginata var．liukiuensis 2n=32，属3C型。讨论了种间在核型上的差异、属的基数、   核型演化趋势和性染色体等问题。  相似文献   

Identification and characterization of a novel 2R,3R-Butanediol dehydrogenase from Bacillus sp. DL01     

《Electronic Journal of Biotechnology》2021

Background2R,3R-butanediol dehydrogenase (R-BDH) and other BDHs contribute to metabolism of 3R/3S-Acetoin (3R/3S-AC) and 2,3-butanediol (2,3-BD), which are important bulk chemicals used in different industries. R-BDH is responsible for oxidizing the hydroxyl group at their (R) configuration. Bacillus species is a promising producer of 3R/3S-AC and 2,3-BD. In this study, R-bdh gene encoding R-BDH from Bacillus sp. DL01 was isolated, expressed and identified.ResultsR-BDH exerted reducing activities towards Diacetyl (DA) and 3R/3S-AC using NADH, and oxidizing activities towards 2R,3R-BD and Meso-BD using NAD⁺, while no activity was detected with 2S,3S-BD. The R-BDH showed its activity at a wide range of temperature (25°C to 65°C) and pH (5.0–8.0). The R-BDH activity was increased significantly by Cd²⁺ when DA, 3R/3S-AC, and Meso-BD were used as substrates, while Fe²⁺ enhanced the activity remarkably at 2R,3R-BD oxidation. Kinetic parameters of the R-BDH from Bacillus sp. DL01 showed the lowest K_m, the highest V_max, and the highest K_cat towards the racemic 3R/3S-AC substrate, also displayed low K_m towards 2R,3R-BD and Meso-BD when compared with other reported R-BDHs.ConclusionsThe R-BDH from Bacillus sp. DL01 was characterized as a novel R-BDH with high enantioselectivity for R-configuration. It considered NAD⁺ and Zn²⁺ dependant enzyme, with a significant affinity towards 3R/3S-AC, 2R,3R-BD, and Meso-BD substrates. Thus, R-BDH is providing an approach to regulate the production of 3R/3S-AC or 2,3-BD from Bacillus sp. DL01.How to cite: Elmahmoudy M, Elfeky N, Zhongji P, et al. Identification and characterization of a novel 2R,3R-Butanediol Dehydrogenase from Bacillus sp. DL01. Electron J Biotechnol 2021;49. https://doi.org/10.1016/j.ejbt.2020.11.002  相似文献   

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

Zhuolin Xiang  Hao Wang  Aakanksha Pant  Giorgia Pastorin  Chengkuo Lee 《Biomicrofluidics》2013,7(6)

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.  相似文献   

Separation of DNA by length in rotational flow: Lattice-Boltzmann-based simulations     

Faihan Alfahani  Michael Antonelli  Jennifer Kreft Pearce 《Biomicrofluidics》2015,9(4)

We use a lattice-Boltzmann based Brownian dynamics simulation to investigate the separation of different lengths of DNA through the combination of a trapping force and the microflow created by counter-rotating vortices. We can separate most long DNA molecules from shorter chains that have lengths differing by as little as 30%. The sensitivity of this technique is determined by the flow rate, size of the trapping region, and the trapping strength. We expect that this technique can be used in microfluidic devices to separate long DNA fragments that result from techniques such as restriction enzyme digests of genomic DNA.The development of novel methods for manipulating biopolymers such as DNA is required for the continued advancement of microfluidic devices. Techniques such as restriction enzyme digests for genomic sequencing rely on the detection of DNA that differ in length by sometimes thousands of base pairs.¹ Methods that separate DNA strands with resolutions on the order of kilobase pairs are required to analyze the products of this technique. To gain an insight into possible techniques to separate polymers, it can be helpful to review the methods to separate particles in microfluidic devices. Experimental work has shown how hydrodynamic mechanisms can lead to separation of particles based on size and deformability.² Eddies, microvortices, and hydrodynamic tweezers have been used to trap and sort particles. The mechanism of the trapping and sorting arises from the differences between interactions of the particles with the fluid.^2–8 In particular, counter-rotating vortices have been used to sort particles and manipulate biopolymers. They have been used to deposit DNA precisely across electrodes⁹ and trap DNA.^10,11 Vortex flow may therefore be a good basis for a technique for sorting DNA by length.Streaming flow has been used in experiments to separate colloids of different size.^3,4 Particles are passed through a channel with a flow field driven by oscillating bubbles and pressure. The flow field becomes a combination of closed and open streamlines. The vortex flow is controlled by the accoustic driving of the bubbles while pressure controls the net flow of the fluid. Larger particles are trapped in the closed vortex flow created by the bubbles, while smaller particles can escape the neighborhood of a bubble in the open streamlines. This leads to efficient separation of particles with size differences as small as 1 μm.Previous work on DNA has shown that counter-rotating vortices can be used to trap DNA dynamically. Long strands of DNA have been observed to stretch between the centers of two counter-rotating vortices. The polymer stays trapped in this state for significant amounts of time.¹² In a different experiment, the vortices were used to thermally cycle the polymer and allow replication via the polymerase chain reaction (PCR). The DNA is also trapped against one wall by a thermophoretic force in these experiments.¹⁰ The strength of the trap is controlled by the gradient in temperature created by a focused infrared laser beam.Trapping DNA at one wall by counter-rotating vortices has also been explored in simulation and found to depend on the Peclet number, Pe = u_maxL/D_m, where u_max is the maximum speed of the vortex, L is the box size, and D_m is the diffusion coefficient of one bead in the polymer chain.¹¹ The trapping rate of the DNA was shown to depend on the competition between the flow compressing the DNA into the trap region and the diffusion of the DNA out of the trap. For the work presented here, Pe ≅ 2000, similar to the previous work done with the same simulation.We extend the previous work to investigate if counter-rotating vortices can be used to separate DNA of different lengths. We use the same type of simulation outlined in Refs. ¹¹ and ^13–17, based on the lattice-Boltzmann method. The simulation method has successfully modeled systems as diverse as thermophoresis of DNA,¹⁴ migration of DNA in a microchannel,¹⁶ and translocation of DNA through a micropore.^17,18 Using this method, the fluid is broken into a lattice with size, ΔL, chosen to be 0.5 μm, and is coupled to a worm-like chain model with Brownian dynamics for the polymer.^19,20 The fluid velocity distribution function, n_i(r, t), describes the fraction of fluid particles with a discretized velocity, c_i, at each lattice site.^21–24 A discrete velocity scheme with nineteen different velocities in three dimensions is used. The velocity distributions will evolve according to ni(r+ciΔτ,t+Δτ)=ni(r,t)+Lij[nj(r,t)−njeq(r,t)],(1)where L is a collision operator such that the fluid relaxes to the equilibrium distribution, nieq given by a second-order expansion of the Maxwell-Boltzmann distribution nieq=ρaci[1+(ci·u)/cs2+uu:(cici−cs2I)/(2cs4)],(2)where cs=1/3ΔLΔτ is the speed of sound. Δτ is the time step for the fluid in the simulation, Δτ = 8.8 × 10⁻⁵. The coefficients a^c_i are determined by satisfying a local isotropy condition ∑iaciciαciβciγciδ=cs4(δαβδγδ+δαγδbetaγ+δαδδβγ).(3)To simplify computation, the velocity distributions are transformed into moment space. The density ρ, momentum density j, and momentum flux density Π are some of the hydrodynamic moments of n_i(r, t). The equilibrium conditions for these three moments are given by ρ=∑nieq,(4) j=∑ci·nieq,(5) Π=∑nieq·cici.(6)L has eigenvalues τ0−1,τ1−1,…,τ18−1, which are the characteristic relaxation times of the moments. The Bhatanagar-Gross-Krook model is used to determine L:²⁵ the non-conserved moments have a single relaxation time, τ_s = 1.0. The conserved moments are density and momentum; for these, τ⁻¹= 0. Fluctuations are added to the fluid stress as in the method of Ladd.²⁴ We have also compared simulations with lattice sizes of 1 μm and 0.25 μm and found no significant differences in the results.The DNA used in the simulation is represented by a worm-like chain model parameterized to capture the dynamics of YOYO-stained λ DNA in bulk solution at room temperature.^15,16,26 Long, flexible DNA is modeled since techniques to separate long DNA molecules with kilobase pair resolution are necessary to complete techniques such as genomic level sequencing using restriction enzyme digests.¹ In addition, such DNA is often used in experiment. Its properties are similar to unstained DNA or DNA stained by other methods.²⁷ Each molecule is represented by N_b beads and N_b − 1 springs. A chain composed of N_b − 1 springs will have a contour length of (N_b − 1) × 2.1 μm. The forces acting on each monomer include: an excluded volume force, a non-linear spring force, the viscous drag force, a random force that produces Brownian motion, a repulsive force from the container walls, and an attractive trapping force only at one wall as shown in Fig. ​Fig.11.¹³ The excluded volume interaction between beads i and j located at r_i and r_j is modeled using the following potential: Uijev=12kBTνNks2(34πSs2)exp(−3|ri−rj|24Ss2),(7)where ν=σk3 is the excluded volume parameter with σ_k = 0.105 μm, the length of one Kuhn segment, N_ks = 19.8 is the number of Kuhn segments per spring, and Ss2=Nks/6)σk2 is the characteristic size of the bead. This excluded volume potential reproduces self avoiding walk statistics. The non-linear spring force is based on force-extension curves from experiments and is given by fijS=kBT2σk[(1−|rj−ri|Nksσk)−2+4|rj−ri|nKσk−1]rj−ri|rj−ri|,(8)which applies when N_ks ≫ 1.Open in a separate windowFIG. 1.Simulation set-up. Arrows indicate direction of fluid flow. The region where the trapping force is active is shaded, and its width (X_stick) is shown. The region used to determine the trapping rate is indicated by the area labeled trap region. Figure is not to scale, the trap region and X_stick are smaller than shown.The beads are modeled as freely draining but subject to a drag force given by F_f = ?6πηa(u_p ? u_f).(9)The beads are also subjected to a random forcing term that is drawn from a Gaussian distribution with zero mean and a variance σ_v = 2k_BTζΔt.(10)The random force reproduces Brownian motion. To conserve total momentum, the momentum change imparted to the beads through their interactions with the fluid is balanced by a momentum change in the fluid. The momentum change is distributed to the three closest fluid lattice sites using a linear interpolation scheme based on the proximity of the lattice site to the polymer beads. Through this momentum transfer, hydrodynamic interactions between the beads occur.The beads are repelled from the walls with a force of magnitude Fwall=250kBTσk3(xbead−xwall)2, xbead>(xwall−1),(11)where the repulsion range is 1ΔL. Each monomer will also be attracted to the top wall by a force with magnitude Fstick=KstickkBTσk3(xbead−xwall+10)2, xbead>(xwall−Xstick)(12)and range X_stick (see Fig. ​Fig.1).1). The sticking force is turned off every one out of one hundred time steps of the polymer (1% of the simulation time steps). We vary both X_stick and K_stick to achieve separation of the polymers.In previous experiments, DNA has been trapped against one wall by using thermophoresis,¹⁰ dielectrophoresis,²⁸ and nanoplasmonic tweezers.²⁹ In the case of thermophoresis, the trap strength (K_stick) can be controlled by tuning the intensity of the temperature gradient and the trap extension (X_stick) can be controlled through the area over which the gradient extends. Both of these are set through focusing of the laser used to produce local heating. Similarly, the trap parameters can be controlled when using plasmonic tweezers by controlling the laser beam exciting the nanoplasmonic structures. In dielectrophoresis, the DNA is trapped by an AC electric field and can be controlled by tuning the frequency and amplitude of the field.In this work, the number of polymers, N_p, is 10 unless otherwise noted, and the container size is 25 ΔL × 50 ΔL × 2 ΔL. The time step for the fluid is Δτ = 8.8 × 10⁻⁵ s, and for the polymer is Δt = 3.7 × 10⁻⁶ s. The total simulation time is over 100 chain relaxation times, allowing sufficient independent samples to perform statistical analysis.Two counter-rotating vortices, shown in 1, are produced by introducing external forces to the fluid bound by walls in the x-direction and periodic in the y and z. Two forces of equal magnitude push on the fluid in the upper x region (12ΔL < x < 25ΔL): one in the +y-direction along y = 10ΔL, and one in the –y-direction along y = 40ΔL. Such counter-rotating vortices can be produced in microfluidic channels using acoustically driven bubbles,^3,4,30 local heating,¹⁰ or plasmonic nanostructures.⁵ The flow speed is controlled by very different external mechanisms in each case. We therefore choose a simple model to produce fluid flow that is not specific to one mechanism.The simulations are started using random initial conditions, and therefore, both lengths of polymer are dispersed throughout the channel. Within a few minutes, the steady state configurations pictured in Figs. ​Figs.22 and ​and33 are reached. We define the steady state as when the number of polymer chains in the trap changes by less than one chain (10 beads) per 1000 polymer time steps. Intermittently, some polymers may still escape and re-enter the trap even in the steady state. Three final configurations are possible: Both the lengths of DNA have become trapped, both lengths continue to rotate freely, or the shorter strand has become trapped while the longer rotates freely. Two of these states leave the polymers mixed; in the third, the strands have separated by size.Open in a separate windowFIG. 2.Snapshots at t = 0Δt (left) and t = 2500Δt (right) showing the separation of 15-bead strands (grey) from 10-bead strands (black) of DNA. For these simulations, K_stick = 55 and Y_stick = 0.7ΔL.Open in a separate windowFIG. 3.Snapshots at t = 0Δt and t = 2500Δt showing the separation of 13-bead strands (grey) from 10-bead strands (black) of DNA. For these simulations, K_stick = 55 and Y_stick = 0.7ΔL as in Fig. ​Fig.2.2. Note that one long polymer is trapped, as well as all of the shorter polymers.By tuning the attractive wall force parameters and fluid flow, the separated steady state can be realized. We first set the flow parameters that allow the larger chains to rotate freely at the center of the vortices while the shorter chains rotate closer to the wall. The trap strength, K_stick, and extension, X_stick, are changed until the shorter polymers do not leave the trap. The same parameters were used to separate 10-bead chains from 15-bead and 13-bead chains.As shown in Fig. ​Fig.2,2, we have been able to separate shorter 10-bead chains from longer 15-bead chains. In the steady state, 97% of the rotating polymers were long polymers averaged over twenty simulations initialized with different random starting conditions. For three simulations, one small polymer would intermittently leave the trap region. In two of these simulations, one long polymer became stably trapped in the steady state. In another simulation, one 15-bead chain was intermittently trapped. On average, the trapped polymers were 5% 15-bead chains and 95% 10-bead chains. Again, 97% of the rotating polymers were 15-bead chains.Simulations conducted with 10-bead and 13-bead chains also showed significant separation of the two sizes as can be seen in Fig. ​Fig.3.3. In the steady state, 30% of the trapped polymers are 13-bead chains and 70% are 10-bead chains, averaged over twenty different random initial starting conditions and 1000 polymer time steps. Only 14.8% of the shorter polymers were not trapped, leading to 85.2% of the freely rotating chains being 13-bead chains. This is therefore a viable test to detect the presence of these longer chains.We have also separated 20-bead chains from 10-bead chains with all of the shorter chains trapped and all of the longer chains freely rotating in the steady-state. These results do not change for twenty different random initial starting conditions and 1000 polymer time steps. None of the longer polymers intermittently enter the trap region nor do any of the shorter polymers intermittently escape.The separation is achieved by tuning the trapping force and flow rate. Strong flows will push all the DNA molecules into the trap. The final state is mixed, with both short and long strands trapped. For flows that are too weak, the short molecules are not sufficiently compressed by the flow and therefore do not enter the trap region. The end state is mixed, with all polymers freely rotating. Separation is achieved when the flow rate is tuned so that the short strands are compressed against the channel wall while the long polymers rotate near the center of the vortices. The trap strength must then be set sufficiently high enough to prevent the short strands from being pulled by the hydrodynamic drag force out of the trap.The mechanism of the separation depends on the differences in the steady state configurations of the polymers and chances of a polymer escaping the trap. As shown in Fig. ​Fig.4,4, both longer and shorter chains are pulled into the trap region by the flow. However, the longer chains have a larger chance of a bead escaping into a region of the flow where the fluid velocity is sufficient to pull the entire strand out of the trap. As shown in Ref. ¹¹, the trapping rate depends on diffusion in a polymer depleted region near the trap, in agreement with classical theory which neglects bead-wall interactions. In addition, the theory depends on the single bead diffusion rate and does not take into account the elastic force holding the beads together. Diffusion becomes as significant as convection in the polymer depleted region leading to dependence on the Peclet number. Since longer polymers have more beads; they have more chances of a single bead diffusing through this layer into the region where convection is again more important. Thus, they are pulled out of the trap at a faster rate than the shorter chains.Open in a separate windowFIG. 4.N, number of beads in the trap region, versus time for 15-bead DNA strands (solid line) and 10-bead DNA strands (dashed line). Here, ΔT = 10000Δt. The simulation parameters are the same as in Fig. ​Fig.22.In addition, longer chains have a second trap resulting from the microflow. As shown in Ref. ¹², DNA in counter-rotating vortices can tumble at the center of one vortex or be stretched between the centers of the two vortices. We have observed both these conformations for the longer polymer strand. They are a stable trajectory for the longer polymer that remains outside of the trapping region. As seen in Fig. ​Fig.4,4, few monomers of the longer chains enter the trap region once the steady state has been reached. However, the shorter polymer rotates at a larger radius than the longer polymer as seen in Fig. ​Fig.5.5. The shorter polymers therefore are pushed back into the trap while the longer strands rotate stably outside the trapping region.Open in a separate windowFIG. 5.Trajectories of 15-bead DNA (grey) and 10-bead DNA (black). The position of each monomer is plotted for 100 consecutive time steps. Note that the longer polymers rotate in the center of the channel while the shorter polymers rotate at the edges. Simulation parameters are the same as in Fig. ​Fig.22.This mechanism is similar to the one proposed for the separation of colloids by size in Refs. ³ and ⁴. In that experimental work, the smaller colloidal particles rotated at larger radii. This allowed the smaller beads to be pushed out of the vicinity of the vortices by the streaming flow, while the larger beads continued to circle. However, in our simulations, we have the additional mechanism of separation based on the increased chance of a longer polymer escaping the trap region. This mechanism is important for maintaining the separation. Long polymers initially in the trap region or which diffuse into the trap would not be able to escape without it.We expect that this technique could be used to detect the sizes of DNA fragments on the order of thousands of base pairs. It relies on the flexibility of the molecule and its interaction with the flow. Common lab procedures such as restriction enzyme digests for DNA fingerprinting can produce these long fragments. Current techniques such as gel electrophoresis require significant time to separate the long strands that move more slowly through the matrix. This effect could therefore be a good candidate for developing a microfluidic analysis that is significantly faster than traditional procedures. Our separation occurs in minutes rather than in hours as for gel electrophoresis.As pointed out in Ref. ², hydrodynamic effects have been shown to be important for microfluidic devices for separation. We have demonstrated, in simulation, a novel hydrodynamic mechanism for separating polymers by length. We hope that these promising calculations will inspire experiments to verify these results.  相似文献   

Label-free density difference amplification-based cell sorting     

Jihwan Song  Minsun Song  Taewook Kang  Dongchoul Kim  Luke P. Lee 《Biomicrofluidics》2014,8(6)

The selective cell separation is a critical step in fundamental life sciences, translational medicine, biotechnology, and energy harvesting. Conventional cell separation methods are fluorescent activated cell sorting and magnetic-activated cell sorting based on fluorescent probes and magnetic particles on cell surfaces. Label-free cell separation methods such as Raman-activated cell sorting, electro-physiologically activated cell sorting, dielectric-activated cell sorting, or inertial microfluidic cell sorting are, however, limited when separating cells of the same kind or cells with similar sizes and dielectric properties, as well as similar electrophysiological phenotypes. Here we report a label-free density difference amplification-based cell sorting (dDACS) without using any external optical, magnetic, electrical forces, or fluidic activations. The conceptual microfluidic design consists of an inlet, hydraulic jump cavity, and multiple outlets. Incoming particles experience gravity, buoyancy, and drag forces in the separation chamber. The height and distance that each particle can reach in the chamber are different and depend on its density, thus allowing for the separation of particles into multiple outlets. The separation behavior of the particles, based on the ratio of the channel heights of the inlet and chamber and Reynolds number has been systematically studied. Numerical simulation reveals that the difference between the heights of only lighter particles with densities close to that of water increases with increasing the ratio of the channel heights, while decreasing Reynolds number can amplify the difference in the heights between the particles considered irrespective of their densities.Separating specific cells from heterogeneous or homogeneous mixtures has been considered as a key step in a wide variety of applications ranging from biomedicine to energy harvesting. For example, the separation and sorting of rare circulating tumor cells (CTCs) from whole blood has gained significant importance in the potential diagnosis and treatment of metastatic cancers.^1,2 Similarly, malaria detection relies on the collection of infected red blood cells (RBCs) from whole blood.^3,4 In addition, the selective separation of lipid-rich microalgae from homogeneous mixtures of microalgae is a promising technique in biomass conversion.⁵To date, conventional cell separation can be done by labelling cells with biomolecules to induce differences in physical properties. For instance, in a fluorescence-activated cell sorter (FACS), cells to be separated are labelled with antibodies or aptamers with fluorescent molecules, and then sorted by applying an electrical potential.^6,7 Similarly, magnetic-activated cell sorter (MACS) uses magnetic.^8,9 Alternatively, label-free cell separation methods have exploited inherent differences in the physical properties (e.g., size and dielectric properties) of different kinds of cells. For example, acoustophoresis forces particles larger than a desired size to move into the center of a fluidic channel by using ultrasonic standing waves.^10–12 Inertial microfluidics takes advantage of curved fluidic channels in order to amplify the size differences between particles.^13,14 Mass-dependent separation of particles based on gravity and hydrodynamic flow was also reported.¹⁵ Particles with different dielectric properties can also be sorted by dielectrophoresis which induces the movement of polarizable particles.^16–18The disadvantage of these methods, however, is that they require external forces and labels that may cause unexpected damage to biological cells.^19–21 More importantly, most methods are limited in separating cells of the same kind or cells with similar sizes and dielectric properties.Here we designed a novel, label-free density difference amplification-based cell sorting (dDACS) that allows the separation of particles with the same size and charge by exploiting subtle differences in density without the use of external forces. Figure 1(a) illustrates the proposed microfluidic model and its underlying mechanism. The conceptual microfluidic system consists of an inlet, a separation chamber (hydraulic jump cavity), and multiple outlets. Particles entering through the inlet experience gravity (F_G), buoyancy (F_B), and drag (F_D) forces in the separation chamber. The net force acting on the particles can be described as F = F_G + F_B + F_D.(1)As particles enter the separation chamber (i.e., hydraulic jump cavity), F_D acting on the particles changes its direction along the streamline. The particles experience additional forces in the y direction due to large tangential angle (Fig. 1(b)). For lighter particles, whose densities are close to that of the surrounding water, F_D becomes comparable to F_G (i.e., in the y direction), while the net force for heavier particles is less affected by this additional contribution of F_D due to a large F_G. As a result, the height (H) and distance (D) that each particle can travel are different depending on its density. The difference in the maximum height (ΔH_max) between two particles with different density (ρ_p1 and ρ_p2) can be further approximated as ΔHmax∝(vyp0)2(vyf−vyp0), (ρp1−ρp2) ,(2)where v_yp0 and v_yf represent the velocity of particle and fluid along the y direction at the entrance of hydraulic jump cavity, respectively.Open in a separate windowFIG. 1.Schematic illustration of label-free density difference amplification-based cell sorting (dDACS), which exploits differences in the densities (ρ₁ > ρ₂) of particles with similar diameters (d) and charge. (a) The conceptual microfluidic design consists of an inlet, a separation chamber (hydraulic jump cavity), and multiple outlets. Incoming particles experience gravity (F_G), buoyancy (F_B), and drag (F_D) forces in the separation chamber, and depending on their densities, the height (H) and distance (D) that each particle is able to reach will be different, allowing the particles to be separated into multiple outlets. (b) Possible microfluidic channel configurations for density-based separation: Uniform channel height (left), gradual channel expansion (middle), and hydraulic jump cavity with sudden channel expansion (right). The height difference between particles with different densities can be amplified by the sudden channel expansion compared to the other two cases due to the relatively large tangential angle, θ of F_D. (|θ₁|≪ |θ₂|) (see Fig. S1 in the supplementary material²²).In comparison with the other two cases (Fig. 1(b) uniform channel height and gradual channel expansion), the height difference between the particles with different densities can be amplified by the sudden channel expansion in the hydraulic jump cavity due to relatively large tangential angle (see supplementary material²²). Therefore, the particles can be separated through the multiple outlets, depending on their height and distance.In order to analyze the separation behavior of particles in the chamber according to differences in their densities, H and D are systematically investigated. The numerical simulations are performed using a commercial CFD software (CFX 14.0; ANSYS 14.0; ANSYS, Inc.). Particles with the same density may have different trajectories in the separation chamber depending on their inlet positions (Fig. 2(a)). Prior to this investigation, the maximum height (H_max) and distance (D_max) for each particle are compared by examining H and D of 100 identical particles at different inlet positions since the inlet position of particles could be controlled.²⁰ Fig. 2(b) shows H_max and D_max of particles with respect to density at a fixed Reynolds number (Re = 0.1). Note that Reynolds number is defined as Re = ρ_fv_fD_h/μ, where ρ_f, v_f, D_h, μ are density of fluid, velocity of the fluid, hydraulic diameter of a channel, and dynamic viscosity of the fluid, respectively. The hydraulic diameter in the Reynolds number is determined with the inlet channel. Particle densities in the range of 1.1 to 2.0 g/cm³ are chosen with the increase of 0.1 g/cm³. These values are quite reasonable in that the densities of many microorganisms such as microalgae are typically within this range and their densities can be varied by 0.2 g/m³ depending on their cellular context.²³ The lighter particles travel with a higher H_max, and longer D_max. With the separation chamber, the height difference between particles with densities of 1.1 and 1.2 g/cm³ can be amplified by about 10 times as compared to that in a channel without the chamber, judging from the position where the 1.1 g/cm³ particle reaches its H_max.Open in a separate windowFIG. 2.Microfluidic particle separation with respect to Reynolds number (Re). (a) Trajectories in the separation chamber of a hundred particles with the same density starting from inlet positions chosen arbitrarily in order to investigate the effect of the inlet positions on the maxima of the height (H_max) and distance (D_max) prior to further simulation. (b) Representative trajectories of particles having different densities from 1.1 to 2.0 g/cm³. (c) The maximum height (H_max) of each particle with respect to Re. (d) Representative maximum distance (D_max) of each particle at Re = 0.1. (Left) Streamline of fluid and representative trajectories of particles with densities of 1.1 and 2.0 g/cm³ in the separation chamber at Re = 0.1 (right).In Fig. 2(c), the values for H_max of particles with respect to Reynolds number (Re) are presented. Since in our study, the maximum height (H_max) and distance (D_max) for each particle were compared by examining H and D of 100 identical particles that are randomly distributed in the channel (throughout all figures), there is little variation in H_max and D_max between each simulation. However, the standard deviation between each simulation is quite small and can be negligible. The H_max values particles at Re = 0.5 with densities of 1.1 g/cm³ and 1.2 g/cm³ are 2.21 × 10³μm and 2.17 × 10³μm, respectively. The difference between H_max of different particles, ΔH_max, increases with decreasing Re. For example, ΔH_max between particles with densities of 1.1 and 2.0 g/cm³ becomes 0.26 × 10³μm at Re = 1.0, but increases to 1.38 × 10³μm as Re decreases to 0.1. As Re increases (velocity of fluid increases), the relative velocity in the y direction between the fluid and the particle increases resulting in increasing of F_D in the y direction since the velocity of particle in the y direction is very small at the entrance of the separation chamber. Thus, contribution of F_D becomes comparable to the net force in the y direction. As a result, most of the particles even in the case of heavier ones travel quite similarly with the streamline, and ΔH_max subsequently decreases. On the other hand, as Re decreases, the contribution of F_G becomes dominant due to the decrease of F_D in the y direction. Consequently, the particles start to cross downwards streamlines as the density of the particles increases and H_max gradually decreases. In addition, irrespective of their densities, ΔH_max of the particles increases with decreasing Re.Fig. 2(d) shows D_max with respect to the density of the particles (left). Different densities of particles show different trajectories due to the relative contribution of F_D to the net force in the y direction depending on the particle density (right). At Re = 0.1, D_max of particles with densities of 1.1 cm³ and 1.2 g/cm³ are 2.91 × 10⁴μm and 1.43 × 10⁴μm, respectively. As the density of a particle increases, its D_max dramatically decreases. The difference in D_max between particles with densities of 1.1 and 1.2 g/cm³ is 1.48 × 10⁴μm, and 0.0037 × 10⁴μm for particles with densities of 1.9 and 2.0 g/cm³. The effect of F_D is stronger compared to that of F_G on lighter particles. Thus, lighter particles travel quite similarly with the streamline and finally have a large D_max. On the other hand, heavier particles where effect of F_G is stronger compared to that of F_D cross downwards streamlines and finally have a small D_max.Next, in order to investigate the separation behavior of particles with respect to the geometry of the microfluidic device, the effect of the ratio of the height of the separation chamber (h_c) to the inlet (h_i) on H_max is investigated as shown in Fig. ​Fig.3.3. Interestingly, H_max of particles with density of 1.1 g/cm³ increases from 1.93 × 10³μm to 6.48 × 10³μm while that of particles with density of 1.9 g/cm³ slightly changes from 0.70 × 10³μm to 0.73 × 10³μm as h_c/h_i increases from 5 to 20.Open in a separate windowFIG. 3.Microfluidic particle separation with respect to the ratio of the height of the inlet (h_i) to the separation chamber (h_c).This result can be attributed to two effects: (1) the change in the streamline and (2) the relative contribution of drag force to the net force depending on the density. With increasing h_c/h_i, dramatic increase in H_max for lighter particles is because the streamline for the lighter ones experiences more vertical displacement in the separation chamber and the contribution of F_D to the net force acting on the lighter one is more significant (see Fig. S2 in the supplementary material²²).Based on this approach, we propose a microfluidic device for the selective separation of the lightest particle. Fig. 4(a) shows one unit (with three outlets) of the proposed microfluidic device that can be connected in series. The ratio of channel heights (h_c/h_i) is set to 20, and the particle densities are in the range of 1.1 ∼ 1.5 g/m³. Fig. 4(b) shows the representative separation behavior of the particles. A portion of the lightest particles (1.1 g/cm³) is selectively separated into the upper and middle outlets, while remaining light particles together with four other heavier particles with densities in the range of 1.2 to 1.5 g/cm³ leave through the lowest outlet. With a single operation of this unit, 40% of the lightest particles are recovered. In addition, the yield increases with increasing number of cycles (Fig. 4(c)).Open in a separate windowFIG. 4.(a) One unit of the proposed microfluidic device for the selective separation of the lightest particle based on the simulation results. Particles are separated into two outlets based on differences in both the height and distance travelled stemming from differences in density. (b) Representative separation behavior of particles observed in the device. (c) The yield of the lightest particle (1.1 g/cm³) with the proposed microfluidic device according to the number of cycles (i.e., this unit is assumed to be connected in series).In summary, we have demonstrated a label-free microfluidic system for the separation of particles according to subtle differences in their densities without external forces. Our microfluidic design consists simply of an inlet, a separation chamber, and multiple outlets. When entering the separation chamber, the particles experience an additional drag force in the y direction, amplifying the difference in both the height and the distance that the particles with different densities can travel within the chamber. At a fixed Reynolds number, with increasing particle density, H_max decreases monotonously, and D_max decreases dramatically. On the other hand, as Reynolds number increases, the difference between the heights of particles with different densities is attenuated. In addition, the simulation reveals that increasing the ratio of the channel heights increases the difference between the heights of particles only when their densities are close to that of the surrounding water. Based on this approach, a microfluidic device for the separation of the lightest particles has been proposed. We expect that our density-based separation design can be beneficial to the selective separation of specific microorganisms such as lipid-rich microalgae for energy harvesting application.  相似文献