Lab-on-a-Chip 시스템의 혈류 역학에 대한 검토: 엔지니어링 관점

Review on Blood Flow Dynamics in Lab-on-a-Chip Systems: An Engineering Perspective

  • Bin-Jie Lai
  • Li-Tao Zhu
  • Zhe Chen*
  • Bo Ouyang*
  • , and 
  • Zheng-Hong Luo*

Abstract

다양한 수송 메커니즘 하에서, “LOC(lab-on-a-chip)” 시스템에서 유동 전단 속도 조건과 밀접한 관련이 있는 혈류 역학은 다양한 수송 현상을 초래하는 것으로 밝혀졌습니다.

본 연구는 적혈구의 동적 혈액 점도 및 탄성 거동과 같은 점탄성 특성의 역할을 통해 LOC 시스템의 혈류 패턴을 조사합니다. 모세관 및 전기삼투압의 주요 매개변수를 통해 LOC 시스템의 혈액 수송 현상에 대한 연구는 실험적, 이론적 및 수많은 수치적 접근 방식을 통해 제공됩니다.

전기 삼투압 점탄성 흐름에 의해 유발되는 교란은 특히 향후 연구 기회를 위해 혈액 및 기타 점탄성 유체를 취급하는 LOC 장치의 혼합 및 분리 기능 향상에 논의되고 적용됩니다. 또한, 본 연구는 보다 정확하고 단순화된 혈류 모델에 대한 요구와 전기역학 효과 하에서 점탄성 유체 흐름에 대한 수치 연구에 대한 강조와 같은 LOC 시스템 하에서 혈류 역학의 수치 모델링의 문제를 식별합니다.

전기역학 현상을 연구하는 동안 제타 전위 조건에 대한 보다 실용적인 가정도 강조됩니다. 본 연구는 모세관 및 전기삼투압에 의해 구동되는 미세유체 시스템의 혈류 역학에 대한 포괄적이고 학제적인 관점을 제공하는 것을 목표로 한다.

KEYWORDS: 

1. Introduction

1.1. Microfluidic Flow in Lab-on-a-Chip (LOC) Systems

Over the past several decades, the ability to control and utilize fluid flow patterns at microscales has gained considerable interest across a myriad of scientific and engineering disciplines, leading to growing interest in scientific research of microfluidics. 

(1) Microfluidics, an interdisciplinary field that straddles physics, engineering, and biotechnology, is dedicated to the behavior, precise control, and manipulation of fluids geometrically constrained to a small, typically submillimeter, scale. 

(2) The engineering community has increasingly focused on microfluidics, exploring different driving forces to enhance working fluid transport, with the aim of accurately and efficiently describing, controlling, designing, and applying microfluidic flow principles and transport phenomena, particularly for miniaturized applications. 

(3) This attention has chiefly been fueled by the potential to revolutionize diagnostic and therapeutic techniques in the biomedical and pharmaceutical sectorsUnder various driving forces in microfluidic flows, intriguing transport phenomena have bolstered confidence in sustainable and efficient applications in fields such as pharmaceutical, biochemical, and environmental science. The “lab-on-a-chip” (LOC) system harnesses microfluidic flow to enable fluid processing and the execution of laboratory tasks on a chip-sized scale. LOC systems have played a vital role in the miniaturization of laboratory operations such as mixing, chemical reaction, separation, flow control, and detection on small devices, where a wide variety of fluids is adapted. Biological fluid flow like blood and other viscoelastic fluids are notably studied among the many working fluids commonly utilized by LOC systems, owing to the optimization in small fluid sample volumed, rapid response times, precise control, and easy manipulation of flow patterns offered by the system under various driving forces. 

(4)The driving forces in blood flow can be categorized as passive or active transport mechanisms and, in some cases, both. Under various transport mechanisms, the unique design of microchannels enables different functionalities in driving, mixing, separating, and diagnosing blood and drug delivery in the blood. 

(5) Understanding and manipulating these driving forces are crucial for optimizing the performance of a LOC system. Such knowledge presents the opportunity to achieve higher efficiency and reliability in addressing cellular level challenges in medical diagnostics, forensic studies, cancer detection, and other fundamental research areas, for applications of point-of-care (POC) devices. 

(6)

1.2. Engineering Approach of Microfluidic Transport Phenomena in LOC Systems

Different transport mechanisms exhibit unique properties at submillimeter length scales in microfluidic devices, leading to significant transport phenomena that differ from those of macroscale flows. An in-depth understanding of these unique transport phenomena under microfluidic systems is often required in fluidic mechanics to fully harness the potential functionality of a LOC system to obtain systematically designed and precisely controlled transport of microfluids under their respective driving force. Fluid mechanics is considered a vital component in chemical engineering, enabling the analysis of fluid behaviors in various unit designs, ranging from large-scale reactors to separation units. Transport phenomena in fluid mechanics provide a conceptual framework for analytically and descriptively explaining why and how experimental results and physiological phenomena occur. The Navier–Stokes (N–S) equation, along with other governing equations, is often adapted to accurately describe fluid dynamics by accounting for pressure, surface properties, velocity, and temperature variations over space and time. In addition, limiting factors and nonidealities for these governing equations should be considered to impose corrections for empirical consistency before physical models are assembled for more accurate controls and efficiency. Microfluidic flow systems often deviate from ideal conditions, requiring adjustments to the standard governing equations. These deviations could arise from factors such as viscous effects, surface interactions, and non-Newtonian fluid properties from different microfluid types and geometrical layouts of microchannels. Addressing these nonidealities supports the refining of theoretical models and prediction accuracy for microfluidic flow behaviors.

The analytical calculation of coupled nonlinear governing equations, which describes the material and energy balances of systems under ideal conditions, often requires considerable computational efforts. However, advancements in computation capabilities, cost reduction, and improved accuracy have made numerical simulations using different numerical and modeling methods a powerful tool for effectively solving these complex coupled equations and modeling various transport phenomena. Computational fluid dynamics (CFD) is a numerical technique used to investigate the spatial and temporal distribution of various flow parameters. It serves as a critical approach to provide insights and reasoning for decision-making regarding the optimal designs involving fluid dynamics, even prior to complex physical model prototyping and experimental procedures. The integration of experimental data, theoretical analysis, and reliable numerical simulations from CFD enables systematic variation of analytical parameters through quantitative analysis, where adjustment to delivery of blood flow and other working fluids in LOC systems can be achieved.

Numerical methods such as the Finite-Difference Method (FDM), Finite-Element-Method (FEM), and Finite-Volume Method (FVM) are heavily employed in CFD and offer diverse approaches to achieve discretization of Eulerian flow equations through filling a mesh of the flow domain. A more in-depth review of numerical methods in CFD and its application for blood flow simulation is provided in Section 2.2.2.

1.3. Scope of the Review

In this Review, we explore and characterize the blood flow phenomena within the LOC systems, utilizing both physiological and engineering modeling approaches. Similar approaches will be taken to discuss capillary-driven flow and electric-osmotic flow (EOF) under electrokinetic phenomena as a passive and active transport scheme, respectively, for blood transport in LOC systems. Such an analysis aims to bridge the gap between physical (experimental) and engineering (analytical) perspectives in studying and manipulating blood flow delivery by different driving forces in LOC systems. Moreover, the Review hopes to benefit the interests of not only blood flow control in LOC devices but also the transport of viscoelastic fluids, which are less studied in the literature compared to that of Newtonian fluids, in LOC systems.

Section 2 examines the complex interplay between viscoelastic properties of blood and blood flow patterns under shear flow in LOC systems, while engineering numerical modeling approaches for blood flow are presented for assistance. Sections 3 and 4 look into the theoretical principles, numerical governing equations, and modeling methodologies for capillary driven flow and EOF in LOC systems as well as their impact on blood flow dynamics through the quantification of key parameters of the two driving forces. Section 5 concludes the characterized blood flow transport processes in LOC systems under these two forces. Additionally, prospective areas of research in improving the functionality of LOC devices employing blood and other viscoelastic fluids and potentially justifying mechanisms underlying microfluidic flow patterns outside of LOC systems are presented. Finally, the challenges encountered in the numerical studies of blood flow under LOC systems are acknowledged, paving the way for further research.

2. Blood Flow Phenomena

ARTICLE SECTIONS

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2.1. Physiological Blood Flow Behavior

Blood, an essential physiological fluid in the human body, serves the vital role of transporting oxygen and nutrients throughout the body. Additionally, blood is responsible for suspending various blood cells including erythrocytes (red blood cells or RBCs), leukocytes (white blood cells), and thrombocytes (blood platelets) in a plasma medium.Among the cells mentioned above, red blood cells (RBCs) comprise approximately 40–45% of the volume of healthy blood. 

(7) An RBC possesses an inherent elastic property with a biconcave shape of an average diameter of 8 μm and a thickness of 2 μm. This biconcave shape maximizes the surface-to-volume ratio, allowing RBCs to endure significant distortion while maintaining their functionality. 

(8,9) Additionally, the biconcave shape optimizes gas exchange, facilitating efficient uptake of oxygen due to the increased surface area. The inherent elasticity of RBCs allows them to undergo substantial distortion from their original biconcave shape and exhibits high flexibility, particularly in narrow channels.RBC deformability enables the cell to deform from a biconcave shape to a parachute-like configuration, despite minor differences in RBC shape dynamics under shear flow between initial cell locations. As shown in Figure 1(a), RBCs initiating with different resting shapes and orientations displaying display a similar deformation pattern 

(10) in terms of its shape. Shear flow induces an inward bending of the cell at the rear position of the rim to the final bending position, 

(11) resulting in an alignment toward the same position of the flow direction.

Figure 1. Images of varying deformation of RBCs and different dynamic blood flow behaviors. (a) The deforming shape behavior of RBCs at four different initiating positions under the same experimental conditions of a flow from left to right, (10) (b) RBC aggregation, (13) (c) CFL region. (18) Reproduced with permission from ref (10). Copyright 2011 Elsevier. Reproduced with permission from ref (13). Copyright 2022 The Authors, under the terms of the Creative Commons (CC BY 4.0) License https://creativecommons.org/licenses/by/4.0/. Reproduced with permission from ref (18). Copyright 2019 Elsevier.

The flexible property of RBCs enables them to navigate through narrow capillaries and traverse a complex network of blood vessels. The deformability of RBCs depends on various factors, including the channel geometry, RBC concentration, and the elastic properties of the RBC membrane. 

(12) Both flexibility and deformability are vital in the process of oxygen exchange among blood and tissues throughout the body, allowing cells to flow in vessels even smaller than the original cell size prior to deforming.As RBCs serve as major components in blood, their collective dynamics also hugely affect blood rheology. RBCs exhibit an aggregation phenomenon due to cell to cell interactions, such as adhesion forces, among populated cells, inducing unique blood flow patterns and rheological behaviors in microfluidic systems. For blood flow in large vessels between a diameter of 1 and 3 cm, where shear rates are not high, a constant viscosity and Newtonian behavior for blood can be assumed. However, under low shear rate conditions (0.1 s

–1) in smaller vessels such as the arteries and venules, which are within a diameter of 0.2 mm to 1 cm, blood exhibits non-Newtonian properties, such as shear-thinning viscosity and viscoelasticity due to RBC aggregation and deformability. The nonlinear viscoelastic property of blood gives rise to a complex relationship between viscosity and shear rate, primarily influenced by the highly elastic behavior of RBCs. A wide range of research on the transient behavior of the RBC shape and aggregation characteristics under varied flow circumstances has been conducted, aiming to obtain a better understanding of the interaction between blood flow shear forces from confined flows.

For a better understanding of the unique blood flow structures and rheological behaviors in microfluidic systems, some blood flow patterns are introduced in the following section.

2.1.1. RBC Aggregation

RBC aggregation is a vital phenomenon to be considered when designing LOC devices due to its impact on the viscosity of the bulk flow. Under conditions of low shear rate, such as in stagnant or low flow rate regions, RBCs tend to aggregate, forming structures known as rouleaux, resembling stacks of coins as shown in Figure 1(b). 

(13) The aggregation of RBCs increases the viscosity at the aggregated region, 

(14) hence slowing down the overall blood flow. However, when exposed to high shear rates, RBC aggregates disaggregate. As shear rates continue to increase, RBCs tend to deform, elongating and aligning themselves with the direction of the flow. 

(15) Such a dynamic shift in behavior from the cells in response to the shear rate forms the basis of the viscoelastic properties observed in whole blood. In essence, the viscosity of the blood varies according to the shear rate conditions, which are related to the velocity gradient of the system. It is significant to take the intricate relationship between shear rate conditions and the change of blood viscosity due to RBC aggregation into account since various flow driving conditions may induce varied effects on the degree of aggregation.

2.1.2. Fåhræus-Lindqvist Effect

The Fåhræus–Lindqvist (FL) effect describes the gradual decrease in the apparent viscosity of blood as the channel diameter decreases. 

(16) This effect is attributed to the migration of RBCs toward the central region in the microchannel, where the flow rate is higher, due to the presence of higher pressure and asymmetric distribution of shear forces. This migration of RBCs, typically observed at blood vessels less than 0.3 mm, toward the higher flow rate region contributes to the change in blood viscosity, which becomes dependent on the channel size. Simultaneously, the increase of the RBC concentration in the central region of the microchannel results in the formation of a less viscous region close to the microchannel wall. This region called the Cell-Free Layer (CFL), is primarily composed of plasma. 

(17) The combination of the FL effect and the following CFL formation provides a unique phenomenon that is often utilized in passive and active plasma separation mechanisms, involving branched and constriction channels for various applications in plasma separation using microfluidic systems.

2.1.3. Cell-Free Layer Formation

In microfluidic blood flow, RBCs form aggregates at the microchannel core and result in a region that is mostly devoid of RBCs near the microchannel walls, as shown in Figure 1(c). 

(18) The region is known as the cell-free layer (CFL). The CFL region is often known to possess a lower viscosity compared to other regions within the blood flow due to the lower viscosity value of plasma when compared to that of the aggregated RBCs. Therefore, a thicker CFL region composed of plasma correlates to a reduced apparent whole blood viscosity. 

(19) A thicker CFL region is often established following the RBC aggregation at the microchannel core under conditions of decreasing the tube diameter. Apart from the dependence on the RBC concentration in the microchannel core, the CFL thickness is also affected by the volume concentration of RBCs, or hematocrit, in whole blood, as well as the deformability of RBCs. Given the influence CFL thickness has on blood flow rheological parameters such as blood flow rate, which is strongly dependent on whole blood viscosity, investigating CFL thickness under shear flow is crucial for LOC systems accounting for blood flow.

2.1.4. Plasma Skimming in Bifurcation Networks

The uneven arrangement of RBCs in bifurcating microchannels, commonly termed skimming bifurcation, arises from the axial migration of RBCs within flowing streams. This uneven distribution contributes to variations in viscosity across differing sizes of bifurcating channels but offers a stabilizing effect. Notably, higher flow rates in microchannels are associated with increased hematocrit levels, resulting in higher viscosity compared with those with lower flow rates. Parametric investigations on bifurcation angle, 

(20) thickness of the CFL, 

(21) and RBC dynamics, including aggregation and deformation, 

(22) may alter the varying viscosity of blood and its flow behavior within microchannels.

2.2. Modeling on Blood Flow Dynamics

2.2.1. Blood Properties and Mathematical Models of Blood Rheology

Under different shear rate conditions in blood flow, the elastic characteristics and dynamic changes of the RBC induce a complex velocity and stress relationship, resulting in the incompatibility of blood flow characterization through standard presumptions of constant viscosity used for Newtonian fluid flow. Blood flow is categorized as a viscoelastic non-Newtonian fluid flow where constitutive equations governing this type of flow take into consideration the nonlinear viscometric properties of blood. To mathematically characterize the evolving blood viscosity and the relationship between the elasticity of RBC and the shear blood flow, respectively, across space and time of the system, a stress tensor (τ) defined by constitutive models is often coupled in the Navier–Stokes equation to account for the collective impact of the constant dynamic viscosity (η) and the elasticity from RBCs on blood flow.The dynamic viscosity of blood is heavily dependent on the shear stress applied to the cell and various parameters from the blood such as hematocrit value, plasma viscosity, mechanical properties of the RBC membrane, and red blood cell aggregation rate. The apparent blood viscosity is considered convenient for the characterization of the relationship between the evolving blood viscosity and shear rate, which can be defined by Casson’s law, as shown in eq 1.

𝜇=𝜏0𝛾˙+2𝜂𝜏0𝛾˙⎯⎯⎯⎯⎯⎯⎯√+𝜂�=�0�˙+2��0�˙+�

(1)where τ

0 is the yield stress–stress required to initiate blood flow motion, η is the Casson rheological constant, and γ̇ is the shear rate. The value of Casson’s law parameters under blood with normal hematocrit level can be defined as τ

0 = 0.0056 Pa and η = 0.0035 Pa·s. 

(23) With the known property of blood and Casson’s law parameters, an approximation can be made to the dynamic viscosity under various flow condition domains. The Power Law model is often employed to characterize the dynamic viscosity in relation to the shear rate, since precise solutions exist for specific geometries and flow circumstances, acting as a fundamental standard for definition. The Carreau and Carreau–Yasuda models can be advantageous over the Power Law model due to their ability to evaluate the dynamic viscosity at low to zero shear rate conditions. However, none of the above-mentioned models consider the memory or other elastic behavior of blood and its RBCs. Some other commonly used mathematical models and their constants for the non-Newtonian viscosity property characterization of blood are listed in Table 1 below. 

(24−26)Table 1. Comparison of Various Non-Newtonian Models for Blood Viscosity 

(24−26)

ModelNon-Newtonian ViscosityParameters
Power Law(2)n = 0.61, k = 0.42
Carreau(3)μ0 = 0.056 Pa·s, μ = 0.00345 Pa·s, λ = 3.1736 s, m = 2.406, a = 0.254
Walburn–Schneck(4)C1 = 0.000797 Pa·s, C2 = 0.0608 Pa·s, C3 = 0.00499, C4 = 14.585 g–1, TPMA = 25 g/L
Carreau–Yasuda(5)μ0 = 0.056 Pa·s, μ = 0.00345 Pa·s, λ = 1.902 s, n = 0.22, a = 1.25
Quemada(6)μp = 0.0012 Pa·s, k = 2.07, k0 = 4.33, γ̇c = 1.88 s–1

The blood rheology is commonly known to be influenced by two key physiological factors, namely, the hematocrit value (H

t) and the fibrinogen concentration (c

f), with an average value of 42% and 0.252 gd·L

–1, respectively. Particularly in low shear conditions, the presence of varying fibrinogen concentrations affects the tendency for aggregation and rouleaux formation, while the occurrence of aggregation is contingent upon specific levels of hematocrit. 

(27) The study from Apostolidis et al. 

(28) modifies the Casson model through emphasizing its reliance on hematocrit and fibrinogen concentration parameter values, owing to the extensive knowledge of the two physiological blood parameters.The viscoelastic response of blood is heavily dependent on the elasticity of the RBC, which is defined by the relationship between the deformation and stress relaxation from RBCs under a specific location of shear flow as a function of the velocity field. The stress tensor is usually characterized by constitutive equations such as the Upper-Convected Maxwell Model 

(29) and the Oldroyd-B model 

(30) to track the molecule effects under shear from different driving forces. The prominent non-Newtonian features, such as shear thinning and yield stress, have played a vital role in the characterization of blood rheology, particularly with respect to the evaluation of yield stress under low shear conditions. The nature of stress measurement in blood, typically on the order of 1 mPa, is challenging due to its low magnitude. The occurrence of the CFL complicates the measurement further due to the significant decrease in apparent viscosity near the wall over time and a consequential disparity in viscosity compared to the bulk region.In addition to shear thinning viscosity and yield stress, the formation of aggregation (rouleaux) from RBCs under low shear rates also contributes to the viscoelasticity under transient flow 

(31) and thixotropy 

(32) of whole blood. Given the difficulty in evaluating viscoelastic behavior of blood under low strain magnitudes and limitations in generalized Newtonian models, the utilization of viscoelastic models is advocated to encompass elasticity and delineate non-shear components within the stress tensor. Extending from the Oldroyd-B model, Anand et al. 

(33) developed a viscoelastic model framework for adapting elasticity within blood samples and predicting non-shear stress components. However, to also address the thixotropic effects, the model developed by Horner et al. 

(34) serves as a more comprehensive approach than the viscoelastic model from Anand et al. Thixotropy 

(32) typically occurs from the structural change of the rouleaux, where low shear rate conditions induce rouleaux formation. Correspondingly, elasticity increases, while elasticity is more representative of the isolated RBCs, under high shear rate conditions. The model of Horner et al. 

(34) considers the contribution of rouleaux to shear stress, taking into account factors such as the characteristic time for Brownian aggregation, shear-induced aggregation, and shear-induced breakage. Subsequent advancements in the model from Horner et al. often revolve around refining the three aforementioned key terms for a more substantial characterization of rouleaux dynamics. Notably, this has led to the recently developed mHAWB model 

(35) and other model iterations to enhance the accuracy of elastic and viscoelastic contributions to blood rheology, including the recently improved model suggested by Armstrong et al. 

(36)

2.2.2. Numerical Methods (FDM, FEM, FVM)

Numerical simulation has become increasingly more significant in analyzing the geometry, boundary layers of flow, and nonlinearity of hyperbolic viscoelastic flow constitutive equations. CFD is a powerful and efficient tool utilizing numerical methods to solve the governing hydrodynamic equations, such as the Navier–Stokes (N–S) equation, continuity equation, and energy conservation equation, for qualitative evaluation of fluid motion dynamics under different parameters. CFD overcomes the challenge of analytically solving nonlinear forms of differential equations by employing numerical methods such as the Finite-Difference Method (FDM), Finite-Element Method (FEM), and Finite-Volume Method (FVM) to discretize and solve the partial differential equations (PDEs), allowing for qualitative reproduction of transport phenomena and experimental observations. Different numerical methods are chosen to cope with various transport systems for optimization of the accuracy of the result and control of error during the discretization process.FDM is a straightforward approach to discretizing PDEs, replacing the continuum representation of equations with a set of finite-difference equations, which is typically applied to structured grids for efficient implementation in CFD programs. 

(37) However, FDM is often limited to simple geometries such as rectangular or block-shaped geometries and struggles with curved boundaries. In contrast, FEM divides the fluid domain into small finite grids or elements, approximating PDEs through a local description of physics. 

(38) All elements contribute to a large, sparse matrix solver. However, FEM may not always provide accurate results for systems involving significant deformation and aggregation of particles like RBCs due to large distortion of grids. 

(39) FVM evaluates PDEs following the conservation laws and discretizes the selected flow domain into small but finite size control volumes, with each grid at the center of a finite volume. 

(40) The divergence theorem allows the conversion of volume integrals of PDEs with divergence terms into surface integrals of surface fluxes across cell boundaries. Due to its conservation property, FVM offers efficient outcomes when dealing with PDEs that embody mass, momentum, and energy conservation principles. Furthermore, widely accessible software packages like the OpenFOAM toolbox 

(41) include a viscoelastic solver, making it an attractive option for viscoelastic fluid flow modeling. 

(42)

2.2.3. Modeling Methods of Blood Flow Dynamics

The complexity in the blood flow simulation arises from deformability and aggregation that RBCs exhibit during their interaction with neighboring cells under different shear rate conditions induced by blood flow. Numerical models coupled with simulation programs have been applied as a groundbreaking method to predict such unique rheological behavior exhibited by RBCs and whole blood. The conventional approach of a single-phase flow simulation is often applied to blood flow simulations within large vessels possessing a moderate shear rate. However, such a method assumes the properties of plasma, RBCs and other cellular components to be evenly distributed as average density and viscosity in blood, resulting in the inability to simulate the mechanical dynamics, such as RBC aggregation under high-shear flow field, inherent in RBCs. To accurately describe the asymmetric distribution of RBC and blood flow, multiphase flow simulation, where numerical simulations of blood flows are often modeled as two immiscible phases, RBCs and blood plasma, is proposed. A common assumption is that RBCs exhibit non-Newtonian behavior while the plasma is treated as a continuous Newtonian phase.Numerous multiphase numerical models have been proposed to simulate the influence of RBCs on blood flow dynamics by different assumptions. In large-scale simulations (above the millimeter range), continuum-based methods are wildly used due to their lower computational demands. 

(43) Eulerian multiphase flow simulations offer the solution of a set of conservation equations for each separate phase and couple the phases through common pressure and interphase exchange coefficients. Xu et al. 

(44) utilized the combined finite-discrete element method (FDEM) to replicate the dynamic behavior and distortion of RBCs subjected to fluidic forces, utilizing the Johnson–Kendall–Roberts model 

(45) to define the adhesive forces of cell-to-cell interactions. The iterative direct-forcing immersed boundary method (IBM) is commonly employed in simulations of the fluid–cell interface of blood. This method effectively captures the intricacies of the thin and flexible RBC membranes within various external flow fields. 

(46) The study by Xu et al. 

(44) also adopts this approach to bridge the fluid dynamics and RBC deformation through IBM. Yoon and You utilized the Maxwell model to define the viscosity of the RBC membrane. 

(47) It was discovered that the Maxwell model could represent the stress relaxation and unloading processes of the cell. Furthermore, the reduced flexibility of an RBC under particular situations such as infection is specified, which was unattainable by the Kelvin–Voigt model 

(48) when compared to the Maxwell model in the literature. The Yeoh hyperplastic material model was also adapted to predict the nonlinear elasticity property of RBCs with FEM employed to discretize the RBC membrane using shell-type elements. Gracka et al. 

(49) developed a numerical CFD model with a finite-volume parallel solver for multiphase blood flow simulation, where an updated Maxwell viscoelasticity model and a Discrete Phase Model are adopted. In the study, the adapted IBM, based on unstructured grids, simulates the flow behavior and shape change of the RBCs through fluid-structure coupling. It was found that the hybrid Euler–Lagrange (E–L) approach 

(50) for the development of the multiphase model offered better results in the simulated CFL region in the microchannels.To study the dynamics of individual behaviors of RBCs and the consequent non-Newtonian blood flow, cell-shape-resolved computational models are often adapted. The use of the boundary integral method has become prevalent in minimizing computational expenses, particularly in the exclusive determination of fluid velocity on the surfaces of RBCs, incorporating the option of employing IBM or particle-based techniques. The cell-shaped-resolved method has enabled an examination of cell to cell interactions within complex ambient or pulsatile flow conditions 

(51) surrounding RBC membranes. Recently, Rydquist et al. 

(52) have looked to integrate statistical information from macroscale simulations to obtain a comprehensive overview of RBC behavior within the immediate proximity of the flow through introduction of respective models characterizing membrane shape definition, tension, bending stresses of RBC membranes.At a macroscopic scale, continuum models have conventionally been adapted for assessing blood flow dynamics through the application of elasticity theory and fluid dynamics. However, particle-based methods are known for their simplicity and adaptability in modeling complex multiscale fluid structures. Meshless methods, such as the boundary element method (BEM), smoothed particle hydrodynamics (SPH), and dissipative particle dynamics (DPD), are often used in particle-based characterization of RBCs and the surrounding fluid. By representing the fluid as discrete particles, meshless methods provide insights into the status and movement of the multiphase fluid. These methods allow for the investigation of cellular structures and microscopic interactions that affect blood rheology. Non-confronting mesh methods like IBM can also be used to couple a fluid solver such as FEM, FVM, or the Lattice Boltzmann Method (LBM) through membrane representation of RBCs. In comparison to conventional CFD methods, LBM has been viewed as a favorable numerical approach for solving the N–S equations and the simulation of multiphase flows. LBM exhibits the notable advantage of being amenable to high-performance parallel computing environments due to its inherently local dynamics. In contrast to DPD and SPH where RBC membranes are modeled as physically interconnected particles, LBM employs the IBM to account for the deformation dynamics of RBCs 

(53,54) under shear flows in complex channel geometries. 

(54,55) However, it is essential to acknowledge that the utilization of LBM in simulating RBC flows often entails a significant computational overhead, being a primary challenge in this context. Krüger et al. 

(56) proposed utilizing LBM as a fluid solver, IBM to couple the fluid and FEM to compute the response of membranes to deformation under immersed fluids. This approach decouples the fluid and membranes but necessitates significant computational effort due to the requirements of both meshes and particles.Despite the accuracy of current blood flow models, simulating complex conditions remains challenging because of the high computational load and cost. Balachandran Nair et al. 

(57) suggested a reduced order model of RBC under the framework of DEM, where the RBC is represented by overlapping constituent rigid spheres. The Morse potential force is adapted to account for the RBC aggregation exhibited by cell to cell interactions among RBCs at different distances. Based upon the IBM, the reduced-order RBC model is adapted to simulate blood flow transport for validation under both single and multiple RBCs with a resolved CFD-DEM solver. 

(58) In the resolved CFD-DEM model, particle sizes are larger than the grid size for a more accurate computation of the surrounding flow field. A continuous forcing approach is taken to describe the momentum source of the governing equation prior to discretization, which is different from a Direct Forcing Method (DFM). 

(59) As no body-conforming moving mesh is required, the continuous forcing approach offers lower complexity and reduced cost when compared to the DFM. Piquet et al. 

(60) highlighted the high complexity of the DFM due to its reliance on calculating an additional immersed boundary flux for the velocity field to ensure its divergence-free condition.The fluid–structure interaction (FSI) method has been advocated to connect the dynamic interplay of RBC membranes and fluid plasma within blood flow such as the coupling of continuum–particle interactions. However, such methodology is generally adapted for anatomical configurations such as arteries 

(61,62) and capillaries, 

(63) where both the structural components and the fluid domain undergo substantial deformation due to the moving boundaries. Due to the scope of the Review being blood flow simulation within microchannels of LOC devices without deformable boundaries, the Review of the FSI method will not be further carried out.In general, three numerical methods are broadly used: mesh-based, particle-based, and hybrid mesh–particle techniques, based on the spatial scale and the fundamental numerical approach, mesh-based methods tend to neglect the effects of individual particles, assuming a continuum and being efficient in terms of time and cost. However, the particle-based approach highlights more of the microscopic and mesoscopic level, where the influence of individual RBCs is considered. A review from Freund et al. 

(64) addressed the three numerical methodologies and their respective modeling approaches of RBC dynamics. Given the complex mechanics and the diverse levels of study concerning numerical simulations of blood and cellular flow, a broad spectrum of numerical methods for blood has been subjected to extensive review. 

(64−70) Ye at al. 

(65) offered an extensive review of the application of the DPD, SPH, and LBM for numerical simulations of RBC, while Rathnayaka et al. 

(67) conducted a review of the particle-based numerical modeling for liquid marbles through drawing parallels to the transport of RBCs in microchannels. A comparative analysis between conventional CFD methods and particle-based approaches for cellular and blood flow dynamic simulation can be found under the review by Arabghahestani et al. 

(66) Literature by Li et al. 

(68) and Beris et al. 

(69) offer an overview of both continuum-based models at micro/macroscales and multiscale particle-based models encompassing various length and temporal dimensions. Furthermore, these reviews deliberate upon the potential of coupling continuum-particle methods for blood plasma and RBC modeling. Arciero et al. 

(70) investigated various modeling approaches encompassing cellular interactions, such as cell to cell or plasma interactions and the individual cellular phases. A concise overview of the reviews is provided in Table 2 for reference.

Table 2. List of Reviews for Numerical Approaches Employed in Blood Flow Simulation

ReferenceNumerical methods
Li et al. (2013) (68)Continuum-based modeling (BIM), particle-based modeling (LBM, LB-FE, SPH, DPD)
Freund (2014) (64)RBC dynamic modeling (continuum-based modeling, complementary discrete microstructure modeling), blood flow dynamic modeling (FDM, IBM, LBM, particle-mesh methods, coupled boundary integral and mesh-based methods, DPD)
Ye et al. (2016) (65)DPD, SPH, LBM, coupled IBM-Smoothed DPD
Arciero et al. (2017) (70)LBM, IBM, DPD, conventional CFD Methods (FDM, FVM, FEM)
Arabghahestani et al. (2019) (66)Particle-based methods (LBM, DPD, direct simulation Monte Carlo, molecular dynamics), SPH, conventional CFD methods (FDM, FVM, FEM)
Beris et al. (2021) (69)DPD, smoothed DPD, IBM, LBM, BIM
Rathnayaka (2022) (67)SPH, CG, LBM

3. Capillary Driven Blood Flow in LOC Systems

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3.1. Capillary Driven Flow Phenomena

Capillary driven (CD) flow is a pivotal mechanism in passive microfluidic flow systems 

(9) such as the blood circulation system and LOC systems. 

(71) CD flow is essentially the movement of a liquid to flow against drag forces, where the capillary effect exerts a force on the liquid at the borders, causing a liquid–air meniscus to flow despite gravity or other drag forces. A capillary pressure drops across the liquid–air interface with surface tension in the capillary radius and contact angle. The capillary effect depends heavily on the interaction between the different properties of surface materials. Different values of contact angles can be manipulated and obtained under varying levels of surface wettability treatments to manipulate the surface properties, resulting in different CD blood delivery rates for medical diagnostic device microchannels. CD flow techniques are appealing for many LOC devices, because they require no external energy. However, due to the passive property of liquid propulsion by capillary forces and the long-term instability of surface treatments on channel walls, the adaptability of CD flow in geometrically complex LOC devices may be limited.

3.2. Theoretical and Numerical Modeling of Capillary Driven Blood Flow

3.2.1. Theoretical Basis and Assumptions of Microfluidic Flow

The study of transport phenomena regarding either blood flow driven by capillary forces or externally applied forces under microfluid systems all demands a comprehensive recognition of the significant differences in flow dynamics between microscale and macroscale. The fundamental assumptions and principles behind fluid transport at the microscale are discussed in this section. Such a comprehension will lay the groundwork for the following analysis of the theoretical basis of capillary forces and their role in blood transport in LOC systems.

At the macroscale, fluid dynamics are often strongly influenced by gravity due to considerable fluid mass. However, the high surface to volume ratio at the microscale shifts the balance toward surface forces (e.g., surface tension and viscous forces), much larger than the inertial force. This difference gives rise to transport phenomena unique to microscale fluid transport, such as the prevalence of laminar flow due to a very low Reynolds number (generally lower than 1). Moreover, the fluid in a microfluidic system is often assumed to be incompressible due to the small flow velocity, indicating constant fluid density in both space and time.Microfluidic flow behaviors are governed by the fundamental principles of mass and momentum conservation, which are encapsulated in the continuity equation and the Navier–Stokes (N–S) equation. The continuity equation describes the conservation of mass, while the N–S equation captures the spatial and temporal variations in velocity, pressure, and other physical parameters. Under the assumption of the negligible influence of gravity in microfluidic systems, the continuity equation and the Eulerian representation of the incompressible N–S equation can be expressed as follows:

∇·𝐮⇀=0∇·�⇀=0

(7)

−∇𝑝+𝜇∇2𝐮⇀+∇·𝝉⇀−𝐅⇀=0−∇�+�∇2�⇀+∇·�⇀−�⇀=0

(8)Here, p is the pressure, u is the fluid viscosity, 

𝝉⇀�⇀ represents the stress tensor, and F is the body force exerted by external forces if present.

3.2.2. Theoretical Basis and Modeling of Capillary Force in LOC Systems

The capillary force is often the major driving force to manipulate and transport blood without an externally applied force in LOC systems. Forces induced by the capillary effect impact the free surface of fluids and are represented not directly in the Navier–Stokes equations but through the pressure boundary conditions of the pressure term p. For hydrophilic surfaces, the liquid generally induces a contact angle between 0° and 30°, encouraging the spread and attraction of fluid under a positive cos θ condition. For this condition, the pressure drop becomes positive and generates a spontaneous flow forward. A hydrophobic solid surface repels the fluid, inducing minimal contact. Generally, hydrophobic solids exhibit a contact angle larger than 90°, inducing a negative value of cos θ. Such a value will result in a negative pressure drop and a flow in the opposite direction. The induced contact angle is often utilized to measure the wall exposure of various surface treatments on channel walls where different wettability gradients and surface tension effects for CD flows are established. Contact angles between different interfaces are obtainable through standard values or experimental methods for reference. 

(72)For the characterization of the induced force by the capillary effect, the Young–Laplace (Y–L) equation 

(73) is widely employed. In the equation, the capillary is considered a pressure boundary condition between the two interphases. Through the Y–L equation, the capillary pressure force can be determined, and subsequently, the continuity and momentum balance equations can be solved to obtain the blood filling rate. Kim et al. 

(74) studied the effects of concentration and exposure time of a nonionic surfactant, Silwet L-77, on the performance of a polydimethylsiloxane (PDMS) microchannel in terms of plasma and blood self-separation. The study characterized the capillary pressure force by incorporating the Y–L equation and further evaluated the effects of the changing contact angle due to different levels of applied channel wall surface treatments. The expression of the Y–L equation utilized by Kim et al. 

(74) is as follows:

𝑃=−𝜎(cos𝜃b+cos𝜃tℎ+cos𝜃l+cos𝜃r𝑤)�=−�(cos⁡�b+cos⁡�tℎ+cos⁡�l+cos⁡�r�)

(9)where σ is the surface tension of the liquid and θ

bθ

tθ

l, and θ

r are the contact angle values between the liquid and the bottom, top, left, and right walls, respectively. A numerical simulation through Coventor software is performed to evaluate the dynamic changes in the filling rate within the microchannel. The simulation results for the blood filling rate in the microchannel are expressed at a specific time stamp, shown in Figure 2. The results portray an increasing instantaneous filling rate of blood in the microchannel following the decrease in contact angle induced by a higher concentration of the nonionic surfactant treated to the microchannel wall.

Figure 2. Numerical simulation of filling rate of capillary driven blood flow under various contact angle conditions at a specific timestamp. (74) Reproduced with permission from ref (74). Copyright 2010 Elsevier.

When in contact with hydrophilic or hydrophobic surfaces, blood forms a meniscus with a contact angle due to surface tension. The Lucas–Washburn (L–W) equation 

(75) is one of the pioneering theoretical definitions for the position of the meniscus over time. In addition, the L–W equation provides the possibility for research to obtain the velocity of the blood formed meniscus through the derivation of the meniscus position. The L–W equation 

(75) can be shown below:

𝐿(𝑡)=𝑅𝜎cos(𝜃)𝑡2𝜇⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯√�(�)=��⁡cos(�)�2�

(10)Here L(t) represents the distance of the liquid driven by the capillary forces. However, the generalized L–W equation solely assumes the constant physical properties from a Newtonian fluid rather than considering the non-Newtonian fluid behavior of blood. Cito et al. 

(76) constructed an enhanced version of the L–W equation incorporating the power law to consider the RBC aggregation and the FL effect. The non-Newtonian fluid apparent viscosity under the Power Law model is defined as

𝜇=𝑘·(𝛾˙)𝑛−1�=�·(�˙)�−1

(11)where γ̇ is the strain rate tensor defined as 

𝛾˙=12𝛾˙𝑖𝑗𝛾˙𝑗𝑖⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯√�˙=12�˙���˙��. The stress tensor term τ is computed as τ = μγ̇

ij. The updated L–W equation by Cito 

(76) is expressed as

𝐿(𝑡)=𝑅[(𝑛+13𝑛+1)(𝜎cos(𝜃)𝑅𝑘)1/𝑛𝑡]𝑛/𝑛+1�(�)=�[(�+13�+1)(�⁡cos(�)��)1/��]�/�+1

(12)where k is the flow consistency index and n is the power law index, respectively. The power law index, from the Power Law model, characterizes the extent of the non-Newtonian behavior of blood. Both the consistency and power law index rely on blood properties such as hematocrit, the appearance of the FL effect, the formation of RBC aggregates, etc. The updated L–W equation computes the location and velocity of blood flow caused by capillary forces at specified time points within the LOC devices, taking into account the effects of blood flow characteristics such as RBC aggregation and the FL effect on dynamic blood viscosity.Apart from the blood flow behaviors triggered by inherent blood properties, unique flow conditions driven by capillary forces that are portrayed under different microchannel geometries also hold crucial implications for CD blood delivery. Berthier et al. 

(77) studied the spontaneous Concus–Finn condition, the condition to initiate the spontaneous capillary flow within a V-groove microchannel, as shown in Figure 3(a) both experimentally and numerically. Through experimental studies, the spontaneous Concus–Finn filament development of capillary driven blood flow is observed, as shown in Figure 3(b), while the dynamic development of blood flow is numerically simulated through CFD simulation.

Figure 3. (a) Sketch of the cross-section of Berthier’s V-groove microchannel, (b) experimental view of blood in the V-groove microchannel, (78) (c) illustration of the dynamic change of the extension of filament from FLOW 3D under capillary flow at three increasing time intervals. (78) Reproduced with permission from ref (78). Copyright 2014 Elsevier.

Berthier et al. 

(77) characterized the contact angle needed for the initiation of the capillary driving force at a zero-inlet pressure, through the half-angle (α) of the V-groove geometry layout, and its relation to the Concus–Finn filament as shown below:

𝜃<𝜋2−𝛼sin𝛼1+2(ℎ2/𝑤)sin𝛼<cos𝜃{�<�2−�sin⁡�1+2(ℎ2/�)⁡sin⁡�<cos⁡�

(13)Three possible regimes were concluded based on the contact angle value for the initiation of flow and development of Concus–Finn filament:

𝜃>𝜃1𝜃1>𝜃>𝜃0𝜃0no SCFSCF without a Concus−Finn filamentSCF without a Concus−Finn filament{�>�1no SCF�1>�>�0SCF without a Concus−Finn filament�0SCF without a Concus−Finn filament

(14)Under Newton’s Law, the force balance with low Reynolds and Capillary numbers results in the neglect of inertial terms. The force balance between the capillary forces and the viscous force induced by the channel wall is proposed to derive the analytical fluid velocity. This relation between the two forces offers insights into the average flow velocity and the penetration distance function dependent on time. The apparent blood viscosity is defined by Berthier et al. 

(78) through Casson’s law, 

(23) given in eq 1. The research used the FLOW-3D program from Flow Science Inc. software, which solves transient, free-surface problems using the FDM in multiple dimensions. The Volume of Fluid (VOF) method 

(79) is utilized to locate and track the dynamic extension of filament throughout the advancing interface within the channel ahead of the main flow at three progressing time stamps, as depicted in Figure 3(c).

4. Electro-osmotic Flow (EOF) in LOC Systems

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The utilization of external forces, such as electric fields, has significantly broadened the possibility of manipulating microfluidic flow in LOC systems. 

(80) Externally applied electric field forces induce a fluid flow from the movement of ions in fluid terms as the “electro-osmotic flow” (EOF).Unique transport phenomena, such as enhanced flow velocity and flow instability, induced by non-Newtonian fluids, particularly viscoelastic fluids, under EOF, have sparked considerable interest in microfluidic devices with simple or complicated geometries within channels. 

(81) However, compared to the study of Newtonian fluids and even other electro-osmotic viscoelastic fluid flows, the literature focusing on the theoretical and numerical modeling of electro-osmotic blood flow is limited due to the complexity of blood properties. Consequently, to obtain a more comprehensive understanding of the complex blood flow behavior under EOF, theoretical and numerical studies of the transport phenomena in the EOF section will be based on the studies of different viscoelastic fluids under EOF rather than that of blood specifically. Despite this limitation, we believe these studies offer valuable insights that can help understand the complex behavior of blood flow under EOF.

4.1. EOF Phenomena

Electro-osmotic flow occurs at the interface between the microchannel wall and bulk phase solution. When in contact with the bulk phase, solution ions are absorbed or dissociated at the solid–liquid interface, resulting in the formation of a charge layer, as shown in Figure 4. This charged channel surface wall interacts with both negative and positive ions in the bulk sample, causing repulsion and attraction forces to create a thin layer of immobilized counterions, known as the Stern layer. The induced electric potential from the wall gradually decreases with an increase in the distance from the wall. The Stern layer potential, commonly termed the zeta potential, controls the intensity of the electrostatic interactions between mobile counterions and, consequently, the drag force from the applied electric field. Next to the Stern layer is the diffuse mobile layer, mainly composed of a mobile counterion. These two layers constitute the “electrical double layer” (EDL), the thickness of which is directly proportional to the ionic strength (concentration) of the bulk fluid. The relationship between the two parameters is characterized by a Debye length (λ

D), expressed as

𝜆𝐷=𝜖𝑘B𝑇2(𝑍𝑒)2𝑐0⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯√��=��B�2(��)2�0

(15)where ϵ is the permittivity of the electrolyte solution, k

B is the Boltzmann constant, T is the electron temperature, Z is the integer valence number, e is the elementary charge, and c

0 is the ionic density.

Figure 4. Schematic diagram of an electro-osmotic flow in a microchannel with negative surface charge. (82) Reproduced with permission from ref (82). Copyright 2012 Woodhead Publishing.

When an electric field is applied perpendicular to the EDL, viscous drag is generated due to the movement of excess ions in the EDL. Electro-osmotic forces can be attributed to the externally applied electric potential (ϕ) and the zeta potential, the system wall induced potential by charged walls (ψ). As illustrated in Figure 4, the majority of ions in the bulk phase have a uniform velocity profile, except for a shear rate condition confined within an extremely thin Stern layer. Therefore, EOF displays a unique characteristic of a “near flat” or plug flow velocity profile, different from the parabolic flow typically induced by pressure-driven microfluidic flow (Hagen–Poiseuille flow). The plug-shaped velocity profile of the EOF possesses a high shear rate above the Stern layer.Overall, the EOF velocity magnitude is typically proportional to the Debye Length (λ

D), zeta potential, and magnitude of the externally applied electric field, while a more viscous liquid reduces the EOF velocity.

4.2. Modeling on Electro-osmotic Viscoelastic Fluid Flow

4.2.1. Theoretical Basis of EOF Mechanisms

The EOF of an incompressible viscoelastic fluid is commonly governed by the continuity and incompressible N–S equations, as shown in eqs 7 and 8, where the stress tensor and the electrostatic force term are coupled. The electro-osmotic body force term F, representing the body force exerted by the externally applied electric force, is defined as 

𝐹⇀=𝑝𝐸𝐸⇀�⇀=���⇀, where ρ

E and 

𝐸⇀�⇀ are the net electric charge density and the applied external electric field, respectively.Numerous models are established to theoretically study the externally applied electric potential and the system wall induced potential by charged walls. The following Laplace equation, expressed as eq 16, is generally adapted and solved to calculate the externally applied potential (ϕ).

∇2𝜙=0∇2�=0

(16)Ion diffusion under applied electric fields, together with mass transport resulting from convection and diffusion, transports ionic solutions in bulk flow under electrokinetic processes. The Nernst–Planck equation can describe these transport methods, including convection, diffusion, and electro-diffusion. Therefore, the Nernst–Planck equation is used to determine the distribution of the ions within the electrolyte. The electric potential induced by the charged channel walls follows the Poisson–Nernst–Plank (PNP) equation, which can be written as eq 17.

∇·[𝐷𝑖∇𝑛𝑖−𝑢⇀𝑛𝑖+𝑛𝑖𝐷𝑖𝑧𝑖𝑒𝑘𝑏𝑇∇(𝜙+𝜓)]=0∇·[��∇��−�⇀��+����������∇(�+�)]=0

(17)where D

in

i, and z

i are the diffusion coefficient, ionic concentration, and ionic valence of the ionic species I, respectively. However, due to the high nonlinearity and numerical stiffness introduced by different lengths and time scales from the PNP equations, the Poisson–Boltzmann (PB) model is often considered the major simplified method of the PNP equation to characterize the potential distribution of the EDL region in microchannels. In the PB model, it is assumed that the ionic species in the fluid follow the Boltzmann distribution. This model is typically valid for steady-state problems where charge transport can be considered negligible, the EDLs do not overlap with each other, and the intrinsic potentials are low. It provides a simplified representation of the potential distribution in the EDL region. The PB equation governing the EDL electric potential distribution is described as

∇2𝜓=(2𝑒𝑧𝑛0𝜀𝜀0)sinh(𝑧𝑒𝜓𝑘b𝑇)∇2�=(2���0��0)⁡sinh(����b�)

(18)where n

0 is the ion bulk concentration, z is the ionic valence, and ε

0 is the electric permittivity in the vacuum. Under low electric potential conditions, an even further simplified model to illustrate the EOF phenomena is the Debye–Hückel (DH) model. The DH model is derived by obtaining a charge density term by expanding the exponential term of the Boltzmann equation in a Taylor series.

4.2.2. EOF Modeling for Viscoelastic Fluids

Many studies through numerical modeling were performed to obtain a deeper understanding of the effect exhibited by externally applied electric fields on viscoelastic flow in microchannels under various geometrical designs. Bello et al. 

(83) found that methylcellulose solution, a non-Newtonian polymer solution, resulted in stronger electro-osmotic mobility in experiments when compared to the predictions by the Helmholtz–Smoluchowski equation, which is commonly used to define the velocity of EOF of a Newtonian fluid. Being one of the pioneers to identify the discrepancies between the EOF of Newtonian and non-Newtonian fluids, Bello et al. attributed such discrepancies to the presence of a very high shear rate in the EDL, resulting in a change in the orientation of the polymer molecules. Park and Lee 

(84) utilized the FVM to solve the PB equation for the characterization of the electric field induced force. In the study, the concept of fractional calculus for the Oldroyd-B model was adapted to illustrate the elastic and memory effects of viscoelastic fluids in a straight microchannel They observed that fluid elasticity and increased ratio of viscoelastic fluid contribution to overall fluid viscosity had a significant impact on the volumetric flow rate and sensitivity of velocity to electric field strength compared to Newtonian fluids. Afonso et al. 

(85) derived an analytical expression for EOF of viscoelastic fluid between parallel plates using the DH model to account for a zeta potential condition below 25 mV. The study established the understanding of the electro-osmotic viscoelastic fluid flow under low zeta potential conditions. Apart from the electrokinetic forces, pressure forces can also be coupled with EOF to generate a unique fluid flow behavior within the microchannel. Sousa et al. 

(86) analytically studied the flow of a standard viscoelastic solution by combining the pressure gradient force with an externally applied electric force. It was found that, at a near wall skimming layer and the outer layer away from the wall, macromolecules migrating away from surface walls in viscoelastic fluids are observed. In the study, the Phan-Thien Tanner (PTT) constitutive model is utilized to characterize the viscoelastic properties of the solution. The approach is found to be valid when the EDL is much thinner than the skimming layer under an enhanced flow rate. Zhao and Yang 

(87) solved the PB equation and Carreau model for the characterization of the EOF mechanism and non-Newtonian fluid respectively through the FEM. The numerical results depict that, different from the EOF of Newtonian fluids, non-Newtonian fluids led to an increase of electro-osmotic mobility for shear thinning fluids but the opposite for shear thickening fluids.Like other fluid transport driving forces, EOF within unique geometrical layouts also portrays unique transport phenomena. Pimenta and Alves 

(88) utilized the FVM to perform numerical simulations of the EOF of viscoelastic fluids considering the PB equation and the Oldroyd-B model, in a cross-slot and flow-focusing microdevices. It was found that electroelastic instabilities are formed due to the development of large stresses inside the EDL with streamlined curvature at geometry corners. Bezerra et al. 

(89) used the FDM to numerically analyze the vortex formation and flow instability from an electro-osmotic non-Newtonian fluid flow in a microchannel with a nozzle geometry and parallel wall geometry setting. The PNP equation is utilized to characterize the charge motion in the EOF and the PTT model for non-Newtonian flow characterization. A constriction geometry is commonly utilized in blood flow adapted in LOC systems due to the change in blood flow behavior under narrow dimensions in a microchannel. Ji et al. 

(90) recently studied the EOF of viscoelastic fluid in a constriction microchannel connected by two relatively big reservoirs on both ends (as seen in Figure 5) filled with the polyacrylamide polymer solution, a viscoelastic fluid, and an incompressible monovalent binary electrolyte solution KCl.

Figure 5. Schematic diagram of a negatively charged constriction microchannel connected to two reservoirs at both ends. An electro-osmotic flow is induced in the system by the induced potential difference between the anode and cathode. (90) Reproduced with permission from ref (90). Copyright 2021 The Authors, under the terms of the Creative Commons (CC BY 4.0) License https://creativecommons.org/licenses/by/4.0/.

In studying the EOF of viscoelastic fluids, the Oldroyd-B model is often utilized to characterize the polymeric stress tensor and the deformation rate of the fluid. The Oldroyd-B model is expressed as follows:

𝜏=𝜂p𝜆(𝐜−𝐈)�=�p�(�−�)

(19)where η

p, λ, c, and I represent the polymer dynamic viscosity, polymer relaxation time, symmetric conformation tensor of the polymer molecules, and the identity matrix, respectively.A log-conformation tensor approach is taken to prevent convergence difficulty induced by the viscoelastic properties. The conformation tensor (c) in the polymeric stress tensor term is redefined by a new tensor (Θ) based on the natural logarithm of the c. The new tensor is defined as

Θ=ln(𝐜)=𝐑ln(𝚲)𝐑Θ=ln(�)=�⁡ln(�)�

(20)in which Λ is the diagonal matrix and R is the orthogonal matrix.Under the new conformation tensor, the induced EOF of a viscoelastic fluid is governed by the continuity and N–S equations adapting the Oldroyd-B model, which is expressed as

∂𝚯∂𝑡+𝐮·∇𝚯=𝛀Θ−ΘΩ+2𝐁+1𝜆(eΘ−𝐈)∂�∂�+�·∇�=�Θ−ΘΩ+2�+1�(eΘ−�)

(21)where Ω and B represent the anti-symmetric matrix and the symmetric traceless matrix of the decomposition of the velocity gradient tensor ∇u, respectively. The conformation tensor can be recovered by c = exp(Θ). The PB model and Laplace equation are utilized to characterize the charged channel wall induced potential and the externally applied potential.The governing equations are numerically solved through the FVM by RheoTool, 

(42) an open-source viscoelastic EOF solver on the OpenFOAM platform. A SIMPLEC (Semi-Implicit Method for Pressure Linked Equations-Consistent) algorithm was applied to solve the velocity-pressure coupling. The pressure field and velocity field were computed by the PCG (Preconditioned Conjugate Gradient) solver and the PBiCG (Preconditioned Biconjugate Gradient) solver, respectively.Ranging magnitudes of an applied electric field or fluid concentration induce both different streamlines and velocity magnitudes at various locations and times of the microchannel. In the study performed by Ji et al., 

(90) notable fluctuation of streamlines and vortex formation is formed at the upper stream entrance of the constriction as shown in Figure 6(a) and (b), respectively, due to the increase of electrokinetic effect, which is seen as a result of the increase in polymeric stress (τ

xx). 

(90) The contraction geometry enhances the EOF velocity within the constriction channel under high E

app condition (600 V/cm). Such phenomena can be attributed to the dependence of electro-osmotic viscoelastic fluid flow on the system wall surface and bulk fluid properties. 

(91)

Figure 6. Schematic diagram of vortex formation and streamlines of EOF depicting flow instability at (a) 1.71 s and (b) 1.75 s. Spatial distribution of the elastic normal stress at (c) high Eapp condition. Streamline of an electro-osmotic flow under Eapp of 600 V/cm (90) for (d) non-Newtonian and (e) Newtonian fluid through a constriction geometry. Reproduced with permission from ref (90). Copyright 2021 The Authors, under the terms of the Creative Commons (CC BY 4.0) License https://creativecommons.org/licenses/by/4.0/.

As elastic normal stress exceeds the local shear stress, flow instability and vortex formation occur. The induced elastic stress under EOF not only enhances the instability of the flow but often generates an irregular secondary flow leading to strong disturbance. 

(92) It is also vital to consider the effect of the constriction layout of microchannels on the alteration of the field strength within the system. The contraction geometry enhances a larger electric field strength compared with other locations of the channel outside the constriction region, resulting in a higher velocity gradient and stronger extension on the polymer within the viscoelastic solution. Following the high shear flow condition, a higher magnitude of stretch for polymer molecules in viscoelastic fluids exhibits larger elastic stresses and enhancement of vortex formation at the region. 

(93)As shown in Figure 6(c), significant elastic normal stress occurs at the inlet of the constriction microchannel. Such occurrence of a polymeric flow can be attributed to the dominating elongational flow, giving rise to high deformation of the polymers within the viscoelastic fluid flow, resulting in higher elastic stress from the polymers. Such phenomena at the entrance result in the difference in velocity streamline as circled in Figure 6(d) compared to that of the Newtonian fluid at the constriction entrance in Figure 6(e). 

(90) The difference between the Newtonian and polymer solution at the exit, as circled in Figure 6(d) and (e), can be attributed to the extrudate swell effect of polymers 

(94) within the viscoelastic fluid flow. The extrudate swell effect illustrates that, as polymers emerge from the constriction exit, they tend to contract in the flow direction and grow in the normal direction, resulting in an extrudate diameter greater than the channel size. The deformation of polymers within the polymeric flow at both the entrance and exit of the contraction channel facilitates the change in shear stress conditions of the flow, leading to the alteration in streamlines of flows for each region.

4.3. EOF Applications in LOC Systems

4.3.1. Mixing in LOC Systems

Rather than relying on the micromixing controlled by molecular diffusion under low Reynolds number conditions, active mixers actively leverage convective instability and vortex formation induced by electro-osmotic flows from alternating current (AC) or direct current (DC) electric fields. Such adaptation is recognized as significant breakthroughs for promotion of fluid mixing in chemical and biological applications such as drug delivery, medical diagnostics, chemical synthesis, and so on. 

(95)Many researchers proposed novel designs of electro-osmosis micromixers coupled with numerical simulations in conjunction with experimental findings to increase their understanding of the role of flow instability and vortex formation in the mixing process under electrokinetic phenomena. Matsubara and Narumi 

(96) numerically modeled the mixing process in a microchannel with four electrodes on each side of the microchannel wall, which generated a disruption through unstable electro-osmotic vortices. It was found that particle mixing was sensitive to both the convection effect induced by the main and secondary vortex within the micromixer and the change in oscillation frequency caused by the supplied AC voltage when the Reynolds number was varied. Qaderi et al. 

(97) adapted the PNP equation to numerically study the effect of the geometry and zeta potential configuration of the microchannel on the mixing process with a combined electro-osmotic pressure driven flow. It was reported that the application of heterogeneous zeta potential configuration enhances the mixing efficiency by around 23% while the height of the hurdles increases the mixing efficiency at most 48.1%. Cho et al. 

(98) utilized the PB model and Laplace equation to numerically simulate the electro-osmotic non-Newtonian fluid mixing process within a wavy and block layout of microchannel walls. The Power Law model is adapted to describe the fluid rheological characteristic. It was found that shear-thinning fluids possess a higher volumetric flow rate, which could result in poorer mixing efficiency compared to that of Newtonian fluids. Numerous studies have revealed that flow instability and vortex generation, in particular secondary vortices produced by barriers or greater magnitudes of heterogeneous zeta potential distribution, enhance mixing by increasing bulk flow velocity and reducing flow distance.To better understand the mechanism of disturbance formed in the system due to externally applied forces, known as electrokinetic instability, literature often utilize the Rayleigh (Ra) number, 

(1) as described below:

𝑅𝑎𝑣=𝑢ev𝑢eo=(𝛾−1𝛾+1)2𝑊𝛿2𝐸el2𝐻2𝜁𝛿Ra�=�ev�eo=(�−1�+1)2��2�el2�2��

(22)where γ is the conductivity ratio of the two streams and can be written as 

𝛾=𝜎el,H𝜎el,L�=�el,H�el,L. The Ra number characterizes the ratio between electroviscous and electro-osmotic flow. A high Ra

v value often results in good mixing. It is evident that fluid properties such as the conductivity (σ) of the two streams play a key role in the formation of disturbances to enhance mixing in microsystems. At the same time, electrokinetic parameters like the zeta potential (ζ) in the Ra number is critical in the characterization of electro-osmotic velocity and a slip boundary condition at the microchannel wall.To understand the mixing result along the channel, the concentration field can be defined and simulated under the assumption of steady state conditions and constant diffusion coefficient for each of the working fluid within the system through the convection–diffusion equation as below:

∂𝑐𝒊∂𝑡+∇⇀(𝑐𝑖𝑢⇀−𝐷𝑖∇⇀𝑐𝒊)=0∂��∂�+∇⇀(���⇀−��∇⇀��)=0

(23)where c

i is the species concentration of species i and D

i is the diffusion coefficient of the corresponding species.The standard deviation of concentration (σ

sd) can be adapted to evaluate the mixing quality of the system. 

(97) The standard deviation for concentration at a specific portion of the channel may be calculated using the equation below:

𝜎sd=∫10(𝐶∗(𝑦∗)−𝐶m)2d𝑦∗∫10d𝑦∗⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯⎯�sd=∫01(�*(�*)−�m)2d�*∫01d�*

(24)where C*(y*) and C

m are the non-dimensional concentration profile and the mean concentration at the portion, respectively. C* is the non-dimensional concentration and can be calculated as 

𝐶∗=𝐶𝐶ref�*=��ref, where C

ref is the reference concentration defined as the bulk solution concentration. The mean concentration profile can be calculated as 

𝐶m=∫10(𝐶∗(𝑦∗)d𝑦∗∫10d𝑦∗�m=∫01(�*(�*)d�*∫01d�*. With the standard deviation of concentration, the mixing efficiency 

(97) can then be calculated as below:

𝜀𝑥=1−𝜎sd𝜎sd,0��=1−�sd�sd,0

(25)where σ

sd,0 is the standard derivation of the case of no mixing. The value of the mixing efficiency is typically utilized in conjunction with the simulated flow field and concentration field to explore the effect of geometrical and electrokinetic parameters on the optimization of the mixing results.

5. Summary

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5.1. Conclusion

Viscoelastic fluids such as blood flow in LOC systems are an essential topic to proceed with diagnostic analysis and research through microdevices in the biomedical and pharmaceutical industries. The complex blood flow behavior is tightly controlled by the viscoelastic characteristics of blood such as the dynamic viscosity and the elastic property of RBCs under various shear rate conditions. Furthermore, the flow behaviors under varied driving forces promote an array of microfluidic transport phenomena that are critical to the management of blood flow and other adapted viscoelastic fluids in LOC systems. This review addressed the blood flow phenomena, the complicated interplay between shear rate and blood flow behaviors, and their numerical modeling under LOC systems through the lens of the viscoelasticity characteristic. Furthermore, a theoretical understanding of capillary forces and externally applied electric forces leads to an in-depth investigation of the relationship between blood flow patterns and the key parameters of the two driving forces, the latter of which is introduced through the lens of viscoelastic fluids, coupling numerical modeling to improve the knowledge of blood flow manipulation in LOC systems. The flow disturbances triggered by the EOF of viscoelastic fluids and their impact on blood flow patterns have been deeply investigated due to their important role and applications in LOC devices. Continuous advancements of various numerical modeling methods with experimental findings through more efficient and less computationally heavy methods have served as an encouraging sign of establishing more accurate illustrations of the mechanisms for multiphase blood and other viscoelastic fluid flow transport phenomena driven by various forces. Such progress is fundamental for the manipulation of unique transport phenomena, such as the generated disturbances, to optimize functionalities offered by microdevices in LOC systems.

The following section will provide further insights into the employment of studied blood transport phenomena to improve the functionality of micro devices adapting LOC technology. A discussion of the novel roles that external driving forces play in microfluidic flow behaviors is also provided. Limitations in the computational modeling of blood flow and electrokinetic phenomena in LOC systems will also be emphasized, which may provide valuable insights for future research endeavors. These discussions aim to provide guidance and opportunities for new paths in the ongoing development of LOC devices that adapt blood flow.

5.2. Future Directions

5.2.1. Electro-osmosis Mixing in LOC Systems

Despite substantial research, mixing results through flow instability and vortex formation phenomena induced by electro-osmotic mixing still deviate from the effective mixing results offered by chaotic mixing results such as those seen in turbulent flows. However, recent discoveries of a mixing phenomenon that is generally observed under turbulent flows are found within electro-osmosis micromixers under low Reynolds number conditions. Zhao 

(99) experimentally discovered a rapid mixing process in an AC applied micromixer, where the power spectrum of concentration under an applied voltage of 20 V

p-p induces a −5/3 slope within a frequency range. This value of the slope is considered as the O–C spectrum in macroflows, which is often visible under relatively high Re conditions, such as the Taylor microscale Reynolds number Re > 500 in turbulent flows. 

(100) However, the Re value in the studied system is less than 1 at the specific location and applied voltage. A secondary flow is also suggested to occur close to microchannel walls, being attributed to the increase of convective instability within the system.Despite the experimental phenomenon proposed by Zhao et al., 

(99) the range of effects induced by vital parameters of an EOF mixing system on the enhanced mixing results and mechanisms of disturbance generated by the turbulent-like flow instability is not further characterized. Such a gap in knowledge may hinder the adaptability and commercialization of the discovery of micromixers. One of the parameters for further evaluation is the conductivity gradient of the fluid flow. A relatively strong conductivity gradient (5000:1) was adopted in the system due to the conductive properties of the two fluids. The high conductivity gradients may contribute to the relatively large Rayleigh number and differences in EDL layer thickness, resulting in an unusual disturbance in laminar flow conditions and enhanced mixing results. However, high conductivity gradients are not always achievable by the working fluids due to diverse fluid properties. The reliance on turbulent-like phenomena and rapid mixing results in a large conductivity gradient should be established to prevent the limited application of fluids for the mixing system. In addition, the proposed system utilizes distinct zeta potential distributions at the top and bottom walls due to their difference in material choices, which may be attributed to the flow instability phenomena. Further studies should be made on varying zeta potential magnitude and distribution to evaluate their effect on the slip boundary conditions of the flow and the large shear rate condition close to the channel wall of EOF. Such a study can potentially offer an optimized condition in zeta potential magnitude through material choices and geometrical layout of the zeta potential for better mixing results and manipulation of mixing fluid dynamics. The two vital parameters mentioned above can be varied with the aid of numerical simulation to understand the effect of parameters on the interaction between electro-osmotic forces and electroviscous forces. At the same time, the relationship of developed streamlines of the simulated velocity and concentration field, following their relationship with the mixing results, under the impact of these key parameters can foster more insight into the range of impact that the two parameters have on the proposed phenomena and the microfluidic dynamic principles of disturbances.

In addition, many of the current investigations of electrokinetic mixers commonly emphasize the fluid dynamics of mixing for Newtonian fluids, while the utilization of biofluids, primarily viscoelastic fluids such as blood, and their distinctive response under shear forces in these novel mixing processes of LOC systems are significantly less studied. To develop more compatible microdevice designs and efficient mixing outcomes for the biomedical industry, it is necessary to fill the knowledge gaps in the literature on electro-osmotic mixing for biofluids, where properties of elasticity, dynamic viscosity, and intricate relationship with shear flow from the fluid are further considered.

5.2.2. Electro-osmosis Separation in LOC Systems

Particle separation in LOC devices, particularly in biological research and diagnostics, is another area where disturbances may play a significant role in optimization. 

(101) Plasma analysis in LOC systems under precise control of blood flow phenomena and blood/plasma separation procedures can detect vital information about infectious diseases from particular antibodies and foreign nucleic acids for medical treatments, diagnostics, and research, 

(102) offering more efficient results and simple operating procedures compared to that of the traditional centrifugation method for blood and plasma separation. However, the adaptability of LOC devices for blood and plasma separation is often hindered by microchannel clogging, where flow velocity and plasma yield from LOC devices is reduced due to occasional RBC migration and aggregation at the filtration entrance of microdevices. 

(103)It is important to note that the EOF induces flow instability close to microchannel walls, which may provide further solutions to clogging for the separation process of the LOC systems. Mohammadi et al. 

(104) offered an anti-clogging effect of RBCs at the blood and plasma separating device filtration entry, adjacent to the surface wall, through RBC disaggregation under high shear rate conditions generated by a forward and reverse EOF direction.

Further theoretical and numerical research can be conducted to characterize the effect of high shear rate conditions near microchannel walls toward the detachment of binding blood cells on surfaces and the reversibility of aggregation. Through numerical modeling with varying electrokinetic parameters to induce different degrees of disturbances or shear conditions at channel walls, it may be possible to optimize and better understand the process of disrupting the forces that bind cells to surface walls and aggregated cells at filtration pores. RBCs that migrate close to microchannel walls are often attracted by the adhesion force between the RBC and the solid surface originating from the van der Waals forces. Following RBC migration and attachment by adhesive forces adjacent to the microchannel walls as shown in Figure 7, the increase in viscosity at the region causes a lower shear condition and encourages RBC aggregation (cell–cell interaction), which clogs filtering pores or microchannels and reduces flow velocity at filtration region. Both the impact that shear forces and disturbances may induce on cell binding forces with surface walls and other cells leading to aggregation may suggest further characterization. Kinetic parameters such as activation energy and the rate-determining step for cell binding composition attachment and detachment should be considered for modeling the dynamics of RBCs and blood flows under external forces in LOC separation devices.

Figure 7. Schematic representations of clogging at a microchannel pore following the sequence of RBC migration, cell attachment to channel walls, and aggregation. (105) Reproduced with permission from ref (105). Copyright 2018 The Authors under the terms of the Creative Commons (CC BY 4.0) License https://creativecommons.org/licenses/by/4.0/.

5.2.3. Relationship between External Forces and Microfluidic Systems

In blood flow, a thicker CFL suggests a lower blood viscosity, suggesting a complex relationship between shear stress and shear rate, affecting the blood viscosity and blood flow. Despite some experimental and numerical studies on electro-osmotic non-Newtonian fluid flow, limited literature has performed an in-depth investigation of the role that applied electric forces and other external forces could play in the process of CFL formation. Additional studies on how shear rates from external forces affect CFL formation and microfluidic flow dynamics can shed light on the mechanism of the contribution induced by external driving forces to the development of a separate phase of layer, similar to CFL, close to the microchannel walls and distinct from the surrounding fluid within the system, then influencing microfluidic flow dynamics.One of the mechanisms of phenomena to be explored is the formation of the Exclusion Zone (EZ) region following a “Self-Induced Flow” (SIF) phenomenon discovered by Li and Pollack, 

(106) as shown in Figure 8(a) and (b), respectively. A spontaneous sustained axial flow is observed when hydrophilic materials are immersed in water, resulting in the buildup of a negative layer of charges, defined as the EZ, after water molecules absorb infrared radiation (IR) energy and break down into H and OH

+.

Figure 8. Schematic representations of (a) the Exclusion Zone region and (b) the Self Induced Flow through visualization of microsphere movement within a microchannel. (106) Reproduced with permission from ref (106). Copyright 2020 The Authors under the terms of the Creative Commons (CC BY 4.0) License https://creativecommons.org/licenses/by/4.0/.

Despite the finding of such a phenomenon, the specific mechanism and role of IR energy have yet to be defined for the process of EZ development. To further develop an understanding of the role of IR energy in such phenomena, a feasible study may be seen through the lens of the relationships between external forces and microfluidic flow. In the phenomena, the increase of SIF velocity under a rise of IR radiation resonant characteristics is shown in the participation of the external electric field near the microchannel walls under electro-osmotic viscoelastic fluid flow systems. The buildup of negative charges at the hydrophilic surfaces in EZ is analogous to the mechanism of electrical double layer formation. Indeed, research has initiated the exploration of the core mechanisms for EZ formation through the lens of the electrokinetic phenomena. 

(107) Such a similarity of the role of IR energy and the transport phenomena of SIF with electrokinetic phenomena paves the way for the definition of the unknown SIF phenomena and EZ formation. Furthermore, Li and Pollack 

(106) suggest whether CFL formation might contribute to a SIF of blood using solely IR radiation, a commonly available source of energy in nature, as an external driving force. The proposition may be proven feasible with the presence of the CFL region next to the negatively charged hydrophilic endothelial glycocalyx layer, coating the luminal side of blood vessels. 

(108) Further research can dive into the resonating characteristics between the formation of the CFL region next to the hydrophilic endothelial glycocalyx layer and that of the EZ formation close to hydrophilic microchannel walls. Indeed, an increase in IR energy is known to rapidly accelerate EZ formation and SIF velocity, depicting similarity to the increase in the magnitude of electric field forces and greater shear rates at microchannel walls affecting CFL formation and EOF velocity. Such correlation depicts a future direction in whether SIF blood flow can be observed and characterized theoretically further through the lens of the relationship between blood flow and shear forces exhibited by external energy.

The intricate link between the CFL and external forces, more specifically the externally applied electric field, can receive further attention to provide a more complete framework for the mechanisms between IR radiation and EZ formation. Such characterization may also contribute to a greater comprehension of the role IR can play in CFL formation next to the endothelial glycocalyx layer as well as its role as a driving force to propel blood flow, similar to the SIF, but without the commonly assumed pressure force from heart contraction as a source of driving force.

5.3. Challenges

Although there have been significant improvements in blood flow modeling under LOC systems over the past decade, there are still notable constraints that may require special attention for numerical simulation applications to benefit the adaptability of the designs and functionalities of LOC devices. Several points that require special attention are mentioned below:

1.The majority of CFD models operate under the relationship between the viscoelasticity of blood and the shear rate conditions of flow. The relative effect exhibited by the presence of highly populated RBCs in whole blood and their forces amongst the cells themselves under complex flows often remains unclearly defined. Furthermore, the full range of cell populations in whole blood requires a much more computational load for numerical modeling. Therefore, a vital goal for future research is to evaluate a reduced modeling method where the impact of cell–cell interaction on the viscoelastic property of blood is considered.
2.Current computational methods on hemodynamics rely on continuum models based upon non-Newtonian rheology at the macroscale rather than at molecular and cellular levels. Careful considerations should be made for the development of a constructive framework for the physical and temporal scales of micro/nanoscale systems to evaluate the intricate relationship between fluid driving forces, dynamic viscosity, and elasticity.
3.Viscoelastic fluids under the impact of externally applied electric forces often deviate from the assumptions of no-slip boundary conditions due to the unique flow conditions induced by externally applied forces. Furthermore, the mechanism of vortex formation and viscoelastic flow instability at laminar flow conditions should be better defined through the lens of the microfluidic flow phenomenon to optimize the prediction of viscoelastic flow across different geometrical layouts. Mathematical models and numerical methods are needed to better predict such disturbance caused by external forces and the viscoelasticity of fluids at such a small scale.
4.Under practical situations, zeta potential distribution at channel walls frequently deviates from the common assumption of a constant distribution because of manufacturing faults or inherent surface charges prior to the introduction of electrokinetic influence. These discrepancies frequently lead to inconsistent surface potential distribution, such as excess positive ions at relatively more negatively charged walls. Accordingly, unpredicted vortex formation and flow instability may occur. Therefore, careful consideration should be given to these discrepancies and how they could trigger the transport process and unexpected results of a microdevice.

Author Information

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  • Corresponding Authors
    • Zhe Chen – Department of Chemical Engineering, School of Chemistry and Chemical Engineering, State Key Laboratory of Metal Matrix Composites, Shanghai Jiao Tong University, Shanghai 200240, P. R. China;  Email: zaccooky@sjtu.edu.cn
    • Bo Ouyang – Department of Chemical Engineering, School of Chemistry and Chemical Engineering, State Key Laboratory of Metal Matrix Composites, Shanghai Jiao Tong University, Shanghai 200240, P. R. China;  Email: bouy93@sjtu.edu.cn
    • Zheng-Hong Luo – Department of Chemical Engineering, School of Chemistry and Chemical Engineering, State Key Laboratory of Metal Matrix Composites, Shanghai Jiao Tong University, Shanghai 200240, P. R. China;  Orcidhttps://orcid.org/0000-0001-9011-6020; Email: luozh@sjtu.edu.cn
  • Authors
    • Bin-Jie Lai – Department of Chemical Engineering, School of Chemistry and Chemical Engineering, State Key Laboratory of Metal Matrix Composites, Shanghai Jiao Tong University, Shanghai 200240, P. R. China;  Orcidhttps://orcid.org/0009-0002-8133-5381
    • Li-Tao Zhu – Department of Chemical Engineering, School of Chemistry and Chemical Engineering, State Key Laboratory of Metal Matrix Composites, Shanghai Jiao Tong University, Shanghai 200240, P. R. China;  Orcidhttps://orcid.org/0000-0001-6514-8864
  • NotesThe authors declare no competing financial interest.

Acknowledgments

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This work was supported by the National Natural Science Foundation of China (No. 22238005) and the Postdoctoral Research Foundation of China (No. GZC20231576).

Vocabulary

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Microfluidicsthe field of technological and scientific study that investigates fluid flow in channels with dimensions between 1 and 1000 μm
Lab-on-a-Chip Technologythe field of research and technological development aimed at integrating the micro/nanofluidic characteristics to conduct laboratory processes on handheld devices
Computational Fluid Dynamics (CFD)the method utilizing computational abilities to predict physical fluid flow behaviors mathematically through solving the governing equations of corresponding fluid flows
Shear Ratethe rate of change in velocity where one layer of fluid moves past the adjacent layer
Viscoelasticitythe property holding both elasticity and viscosity characteristics relying on the magnitude of applied shear stress and time-dependent strain
Electro-osmosisthe flow of fluid under an applied electric field when charged solid surface is in contact with the bulk fluid
Vortexthe rotating motion of a fluid revolving an axis line

References

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This article references 108 other publications.

  1. 1Neethirajan, S.; Kobayashi, I.; Nakajima, M.; Wu, D.; Nandagopal, S.; Lin, F. Microfluidics for food, agriculture and biosystems industries. Lab Chip 201111 (9), 1574– 1586,  DOI: 10.1039/c0lc00230eViewGoogle Scholar
  2. 2Whitesides, G. M. The origins and the future of microfluidics. Nature 2006442 (7101), 368– 373,  DOI: 10.1038/nature05058ViewGoogle Scholar
  3. 3Burklund, A.; Tadimety, A.; Nie, Y.; Hao, N.; Zhang, J. X. J. Chapter One – Advances in diagnostic microfluidics; Elsevier, 2020; DOI:  DOI: 10.1016/bs.acc.2019.08.001 .ViewGoogle Scholar
  4. 4Abdulbari, H. A. Chapter 12 – Lab-on-a-chip for analysis of blood. In Nanotechnology for Hematology, Blood Transfusion, and Artificial Blood; Denizli, A., Nguyen, T. A., Rajan, M., Alam, M. F., Rahman, K., Eds.; Elsevier, 2022; pp 265– 283.ViewGoogle Scholar
  5. 5Vladisavljević, G. T.; Khalid, N.; Neves, M. A.; Kuroiwa, T.; Nakajima, M.; Uemura, K.; Ichikawa, S.; Kobayashi, I. Industrial lab-on-a-chip: Design, applications and scale-up for drug discovery and delivery. Advanced Drug Delivery Reviews 201365 (11), 1626– 1663,  DOI: 10.1016/j.addr.2013.07.017ViewGoogle Scholar
  6. 6Kersaudy-Kerhoas, M.; Dhariwal, R.; Desmulliez, M. P. Y.; Jouvet, L. Hydrodynamic blood plasma separation in microfluidic channels. Microfluid. Nanofluid. 20108 (1), 105– 114,  DOI: 10.1007/s10404-009-0450-5ViewGoogle Scholar
  7. 7Popel, A. S.; Johnson, P. C. Microcirculation and Hemorheology. Annu. Rev. Fluid Mech. 200537 (1), 43– 69,  DOI: 10.1146/annurev.fluid.37.042604.133933ViewGoogle Scholar
  8. 8Fedosov, D. A.; Peltomäki, M.; Gompper, G. Deformation and dynamics of red blood cells in flow through cylindrical microchannels. Soft Matter 201410 (24), 4258– 4267,  DOI: 10.1039/C4SM00248BViewGoogle Scholar
  9. 9Chakraborty, S. Dynamics of capillary flow of blood into a microfluidic channel. Lab Chip 20055 (4), 421– 430,  DOI: 10.1039/b414566fViewGoogle Scholar
  10. 10Tomaiuolo, G.; Guido, S. Start-up shape dynamics of red blood cells in microcapillary flow. Microvascular Research 201182 (1), 35– 41,  DOI: 10.1016/j.mvr.2011.03.004ViewGoogle Scholar
  11. 11Sherwood, J. M.; Dusting, J.; Kaliviotis, E.; Balabani, S. The effect of red blood cell aggregation on velocity and cell-depleted layer characteristics of blood in a bifurcating microchannel. Biomicrofluidics 20126 (2), 24119,  DOI: 10.1063/1.4717755ViewGoogle Scholar
  12. 12Nader, E.; Skinner, S.; Romana, M.; Fort, R.; Lemonne, N.; Guillot, N.; Gauthier, A.; Antoine-Jonville, S.; Renoux, C.; Hardy-Dessources, M.-D. Blood Rheology: Key Parameters, Impact on Blood Flow, Role in Sickle Cell Disease and Effects of Exercise. Frontiers in Physiology 201910, 01329,  DOI: 10.3389/fphys.2019.01329ViewGoogle Scholar
  13. 13Trejo-Soto, C.; Lázaro, G. R.; Pagonabarraga, I.; Hernández-Machado, A. Microfluidics Approach to the Mechanical Properties of Red Blood Cell Membrane and Their Effect on Blood Rheology. Membranes 202212 (2), 217,  DOI: 10.3390/membranes12020217ViewGoogle Scholar
  14. 14Wagner, C.; Steffen, P.; Svetina, S. Aggregation of red blood cells: From rouleaux to clot formation. Comptes Rendus Physique 201314 (6), 459– 469,  DOI: 10.1016/j.crhy.2013.04.004ViewGoogle Scholar
  15. 15Kim, H.; Zhbanov, A.; Yang, S. Microfluidic Systems for Blood and Blood Cell Characterization. Biosensors 202313 (1), 13,  DOI: 10.3390/bios13010013ViewGoogle Scholar
  16. 16Fåhræus, R.; Lindqvist, T. THE VISCOSITY OF THE BLOOD IN NARROW CAPILLARY TUBES. American Journal of Physiology-Legacy Content 193196 (3), 562– 568,  DOI: 10.1152/ajplegacy.1931.96.3.562ViewGoogle Scholar
  17. 17Ascolese, M.; Farina, A.; Fasano, A. The Fåhræus-Lindqvist effect in small blood vessels: how does it help the heart?. J. Biol. Phys. 201945 (4), 379– 394,  DOI: 10.1007/s10867-019-09534-4ViewGoogle Scholar
  18. 18Bento, D.; Fernandes, C. S.; Miranda, J. M.; Lima, R. In vitro blood flow visualizations and cell-free layer (CFL) measurements in a microchannel network. Experimental Thermal and Fluid Science 2019109, 109847,  DOI: 10.1016/j.expthermflusci.2019.109847ViewGoogle Scholar
  19. 19Namgung, B.; Ong, P. K.; Wong, Y. H.; Lim, D.; Chun, K. J.; Kim, S. A comparative study of histogram-based thresholding methods for the determination of cell-free layer width in small blood vessels. Physiological Measurement 201031 (9), N61,  DOI: 10.1088/0967-3334/31/9/N01ViewGoogle Scholar
  20. 20Hymel, S. J.; Lan, H.; Fujioka, H.; Khismatullin, D. B. Cell trapping in Y-junction microchannels: A numerical study of the bifurcation angle effect in inertial microfluidics. Phys. Fluids (1994) 201931 (8), 082003,  DOI: 10.1063/1.5113516ViewGoogle Scholar
  21. 21Li, X.; Popel, A. S.; Karniadakis, G. E. Blood-plasma separation in Y-shaped bifurcating microfluidic channels: a dissipative particle dynamics simulation study. Phys. Biol. 20129 (2), 026010,  DOI: 10.1088/1478-3975/9/2/026010ViewGoogle Scholar
  22. 22Yin, X.; Thomas, T.; Zhang, J. Multiple red blood cell flows through microvascular bifurcations: Cell free layer, cell trajectory, and hematocrit separation. Microvascular Research 201389, 47– 56,  DOI: 10.1016/j.mvr.2013.05.002ViewGoogle Scholar
  23. 23Shibeshi, S. S.; Collins, W. E. The Rheology of Blood Flow in a Branched Arterial System. Appl. Rheol 200515 (6), 398– 405,  DOI: 10.1515/arh-2005-0020ViewGoogle Scholar
  24. 24Sequeira, A.; Janela, J. An Overview of Some Mathematical Models of Blood Rheology. In A Portrait of State-of-the-Art Research at the Technical University of Lisbon; Pereira, M. S., Ed.; Springer Netherlands: Dordrecht, 2007; pp 65– 87.ViewGoogle Scholar
  25. 25Walburn, F. J.; Schneck, D. J. A constitutive equation for whole human blood. Biorheology 197613, 201– 210,  DOI: 10.3233/BIR-1976-13307ViewGoogle Scholar
  26. 26Quemada, D. A rheological model for studying the hematocrit dependence of red cell-red cell and red cell-protein interactions in blood. Biorheology 198118, 501– 516,  DOI: 10.3233/BIR-1981-183-615ViewGoogle Scholar
  27. 27Varchanis, S.; Dimakopoulos, Y.; Wagner, C.; Tsamopoulos, J. How viscoelastic is human blood plasma?. Soft Matter 201814 (21), 4238– 4251,  DOI: 10.1039/C8SM00061AViewGoogle Scholar
  28. 28Apostolidis, A. J.; Moyer, A. P.; Beris, A. N. Non-Newtonian effects in simulations of coronary arterial blood flow. J. Non-Newtonian Fluid Mech. 2016233, 155– 165,  DOI: 10.1016/j.jnnfm.2016.03.008ViewGoogle Scholar
  29. 29Luo, X. Y.; Kuang, Z. B. A study on the constitutive equation of blood. J. Biomech. 199225 (8), 929– 934,  DOI: 10.1016/0021-9290(92)90233-QViewGoogle Scholar
  30. 30Oldroyd, J. G.; Wilson, A. H. On the formulation of rheological equations of state. Proceedings of the Royal Society of London. Series A. Mathematical and Physical Sciences 1950200 (1063), 523– 541,  DOI: 10.1098/rspa.1950.0035ViewGoogle Scholar
  31. 31Prado, G.; Farutin, A.; Misbah, C.; Bureau, L. Viscoelastic transient of confined red blood cells. Biophys J. 2015108 (9), 2126– 2136,  DOI: 10.1016/j.bpj.2015.03.046ViewGoogle Scholar
  32. 32Huang, C. R.; Pan, W. D.; Chen, H. Q.; Copley, A. L. Thixotropic properties of whole blood from healthy human subjects. Biorheology 198724 (6), 795– 801,  DOI: 10.3233/BIR-1987-24630ViewGoogle Scholar
  33. 33Anand, M.; Kwack, J.; Masud, A. A new generalized Oldroyd-B model for blood flow in complex geometries. International Journal of Engineering Science 201372, 78– 88,  DOI: 10.1016/j.ijengsci.2013.06.009ViewGoogle Scholar
  34. 34Horner, J. S.; Armstrong, M. J.; Wagner, N. J.; Beris, A. N. Investigation of blood rheology under steady and unidirectional large amplitude oscillatory shear. J. Rheol. 201862 (2), 577– 591,  DOI: 10.1122/1.5017623ViewGoogle Scholar
  35. 35Horner, J. S.; Armstrong, M. J.; Wagner, N. J.; Beris, A. N. Measurements of human blood viscoelasticity and thixotropy under steady and transient shear and constitutive modeling thereof. J. Rheol. 201963 (5), 799– 813,  DOI: 10.1122/1.5108737ViewGoogle Scholar
  36. 36Armstrong, M.; Tussing, J. A methodology for adding thixotropy to Oldroyd-8 family of viscoelastic models for characterization of human blood. Phys. Fluids 202032 (9), 094111,  DOI: 10.1063/5.0022501ViewGoogle Scholar
  37. 37Crank, J.; Nicolson, P. A practical method for numerical evaluation of solutions of partial differential equations of the heat-conduction type. Mathematical Proceedings of the Cambridge Philosophical Society 194743 (1), 50– 67,  DOI: 10.1017/S0305004100023197ViewGoogle Scholar
  38. 38Clough, R. W. Original formulation of the finite element method. Finite Elements in Analysis and Design 19907 (2), 89– 101,  DOI: 10.1016/0168-874X(90)90001-UViewGoogle Scholar
  39. 39Liu, W. K.; Liu, Y.; Farrell, D.; Zhang, L.; Wang, X. S.; Fukui, Y.; Patankar, N.; Zhang, Y.; Bajaj, C.; Lee, J.Immersed finite element method and its applications to biological systems. Computer Methods in Applied Mechanics and Engineering 2006195 (13), 1722– 1749,  DOI: 10.1016/j.cma.2005.05.049ViewGoogle Scholar
  40. 40Lopes, D.; Agujetas, R.; Puga, H.; Teixeira, J.; Lima, R.; Alejo, J. P.; Ferrera, C. Analysis of finite element and finite volume methods for fluid-structure interaction simulation of blood flow in a real stenosed artery. International Journal of Mechanical Sciences 2021207, 106650,  DOI: 10.1016/j.ijmecsci.2021.106650ViewGoogle Scholar
  41. 41Favero, J. L.; Secchi, A. R.; Cardozo, N. S. M.; Jasak, H. Viscoelastic flow analysis using the software OpenFOAM and differential constitutive equations. J. Non-Newtonian Fluid Mech. 2010165 (23), 1625– 1636,  DOI: 10.1016/j.jnnfm.2010.08.010ViewGoogle Scholar
  42. 42Pimenta, F.; Alves, M. A. Stabilization of an open-source finite-volume solver for viscoelastic fluid flows. J. Non-Newtonian Fluid Mech. 2017239, 85– 104,  DOI: 10.1016/j.jnnfm.2016.12.002ViewGoogle Scholar
  43. 43Chee, C. Y.; Lee, H. P.; Lu, C. Using 3D fluid-structure interaction model to analyse the biomechanical properties of erythrocyte. Phys. Lett. A 2008372 (9), 1357– 1362,  DOI: 10.1016/j.physleta.2007.09.067ViewGoogle Scholar
  44. 44Xu, D.; Kaliviotis, E.; Munjiza, A.; Avital, E.; Ji, C.; Williams, J. Large scale simulation of red blood cell aggregation in shear flows. J. Biomech. 201346 (11), 1810– 1817,  DOI: 10.1016/j.jbiomech.2013.05.010ViewGoogle Scholar
  45. 45Johnson, K. L.; Kendall, K.; Roberts, A. Surface energy and the contact of elastic solids. Proceedings of the royal society of London. A. mathematical and physical sciences 1971324 (1558), 301– 313,  DOI: 10.1098/rspa.1971.0141ViewGoogle Scholar
  46. 46Shi, L.; Pan, T.-W.; Glowinski, R. Deformation of a single red blood cell in bounded Poiseuille flows. Phys. Rev. E 201285 (1), 016307,  DOI: 10.1103/PhysRevE.85.016307ViewGoogle Scholar
  47. 47Yoon, D.; You, D. Continuum modeling of deformation and aggregation of red blood cells. J. Biomech. 201649 (11), 2267– 2279,  DOI: 10.1016/j.jbiomech.2015.11.027ViewGoogle Scholar
  48. 48Mainardi, F.; Spada, G. Creep, relaxation and viscosity properties for basic fractional models in rheology. European Physical Journal Special Topics 2011193 (1), 133– 160,  DOI: 10.1140/epjst/e2011-01387-1ViewGoogle Scholar
  49. 49Gracka, M.; Lima, R.; Miranda, J. M.; Student, S.; Melka, B.; Ostrowski, Z. Red blood cells tracking and cell-free layer formation in a microchannel with hyperbolic contraction: A CFD model validation. Computer Methods and Programs in Biomedicine 2022226, 107117,  DOI: 10.1016/j.cmpb.2022.107117ViewGoogle Scholar
  50. 50Aryan, H.; Beigzadeh, B.; Siavashi, M. Euler-Lagrange numerical simulation of improved magnetic drug delivery in a three-dimensional CT-based carotid artery bifurcation. Computer Methods and Programs in Biomedicine 2022219, 106778,  DOI: 10.1016/j.cmpb.2022.106778ViewGoogle Scholar
  51. 51Czaja, B.; Závodszky, G.; Azizi Tarksalooyeh, V.; Hoekstra, A. G. Cell-resolved blood flow simulations of saccular aneurysms: effects of pulsatility and aspect ratio. J. R Soc. Interface 201815 (146), 20180485,  DOI: 10.1098/rsif.2018.0485ViewGoogle Scholar
  52. 52Rydquist, G.; Esmaily, M. A cell-resolved, Lagrangian solver for modeling red blood cell dynamics in macroscale flows. J. Comput. Phys. 2022461, 111204,  DOI: 10.1016/j.jcp.2022.111204ViewGoogle Scholar
  53. 53Dadvand, A.; Baghalnezhad, M.; Mirzaee, I.; Khoo, B. C.; Ghoreishi, S. An immersed boundary-lattice Boltzmann approach to study the dynamics of elastic membranes in viscous shear flows. Journal of Computational Science 20145 (5), 709– 718,  DOI: 10.1016/j.jocs.2014.06.006ViewGoogle Scholar
  54. 54Krüger, T.; Holmes, D.; Coveney, P. V. Deformability-based red blood cell separation in deterministic lateral displacement devices─A simulation study. Biomicrofluidics 20148 (5), 054114,  DOI: 10.1063/1.4897913ViewGoogle Scholar
  55. 55Takeishi, N.; Ito, H.; Kaneko, M.; Wada, S. Deformation of a Red Blood Cell in a Narrow Rectangular Microchannel. Micromachines 201910 (3), 199,  DOI: 10.3390/mi10030199ViewGoogle Scholar
  56. 56Krüger, T.; Varnik, F.; Raabe, D. Efficient and accurate simulations of deformable particles immersed in a fluid using a combined immersed boundary lattice Boltzmann finite element method. Computers & Mathematics with Applications 201161 (12), 3485– 3505,  DOI: 10.1016/j.camwa.2010.03.057ViewGoogle Scholar
  57. 57Balachandran Nair, A. N.; Pirker, S.; Umundum, T.; Saeedipour, M. A reduced-order model for deformable particles with application in bio-microfluidics. Computational Particle Mechanics 20207 (3), 593– 601,  DOI: 10.1007/s40571-019-00283-8ViewGoogle Scholar
  58. 58Balachandran Nair, A. N.; Pirker, S.; Saeedipour, M. Resolved CFD-DEM simulation of blood flow with a reduced-order RBC model. Computational Particle Mechanics 20229 (4), 759– 774,  DOI: 10.1007/s40571-021-00441-xViewGoogle Scholar
  59. 59Mittal, R.; Iaccarino, G. IMMERSED BOUNDARY METHODS. Annu. Rev. Fluid Mech. 200537 (1), 239– 261,  DOI: 10.1146/annurev.fluid.37.061903.175743ViewGoogle Scholar
  60. 60Piquet, A.; Roussel, O.; Hadjadj, A. A comparative study of Brinkman penalization and direct-forcing immersed boundary methods for compressible viscous flows. Computers & Fluids 2016136, 272– 284,  DOI: 10.1016/j.compfluid.2016.06.001ViewGoogle Scholar
  61. 61Akerkouch, L.; Le, T. B. A Hybrid Continuum-Particle Approach for Fluid-Structure Interaction Simulation of Red Blood Cells in Fluid Flows. Fluids 20216 (4), 139,  DOI: 10.3390/fluids6040139ViewGoogle Scholar
  62. 62Barker, A. T.; Cai, X.-C. Scalable parallel methods for monolithic coupling in fluid-structure interaction with application to blood flow modeling. J. Comput. Phys. 2010229 (3), 642– 659,  DOI: 10.1016/j.jcp.2009.10.001ViewGoogle Scholar
  63. 63Cetin, A.; Sahin, M. A monolithic fluid-structure interaction framework applied to red blood cells. International Journal for Numerical Methods in Biomedical Engineering 201935 (2), e3171  DOI: 10.1002/cnm.3171ViewGoogle Scholar
  64. 64Freund, J. B. Numerical Simulation of Flowing Blood Cells. Annu. Rev. Fluid Mech. 201446 (1), 67– 95,  DOI: 10.1146/annurev-fluid-010313-141349ViewGoogle Scholar
  65. 65Ye, T.; Phan-Thien, N.; Lim, C. T. Particle-based simulations of red blood cells─A review. J. Biomech. 201649 (11), 2255– 2266,  DOI: 10.1016/j.jbiomech.2015.11.050ViewGoogle Scholar
  66. 66Arabghahestani, M.; Poozesh, S.; Akafuah, N. K. Advances in Computational Fluid Mechanics in Cellular Flow Manipulation: A Review. Applied Sciences 20199 (19), 4041,  DOI: 10.3390/app9194041ViewGoogle Scholar
  67. 67Rathnayaka, C. M.; From, C. S.; Geekiyanage, N. M.; Gu, Y. T.; Nguyen, N. T.; Sauret, E. Particle-Based Numerical Modelling of Liquid Marbles: Recent Advances and Future Perspectives. Archives of Computational Methods in Engineering 202229 (5), 3021– 3039,  DOI: 10.1007/s11831-021-09683-7ViewGoogle Scholar
  68. 68Li, X.; Vlahovska, P. M.; Karniadakis, G. E. Continuum- and particle-based modeling of shapes and dynamics of red blood cells in health and disease. Soft Matter 20139 (1), 28– 37,  DOI: 10.1039/C2SM26891DViewGoogle Scholar
  69. 69Beris, A. N.; Horner, J. S.; Jariwala, S.; Armstrong, M. J.; Wagner, N. J. Recent advances in blood rheology: a review. Soft Matter 202117 (47), 10591– 10613,  DOI: 10.1039/D1SM01212FViewGoogle Scholar
  70. 70Arciero, J.; Causin, P.; Malgaroli, F. Mathematical methods for modeling the microcirculation. AIMS Biophysics 20174 (3), 362– 399,  DOI: 10.3934/biophy.2017.3.362ViewGoogle Scholar
  71. 71Maria, M. S.; Chandra, T. S.; Sen, A. K. Capillary flow-driven blood plasma separation and on-chip analyte detection in microfluidic devices. Microfluid. Nanofluid. 201721 (4), 72,  DOI: 10.1007/s10404-017-1907-6ViewGoogle Scholar
  72. 72Huhtamäki, T.; Tian, X.; Korhonen, J. T.; Ras, R. H. A. Surface-wetting characterization using contact-angle measurements. Nat. Protoc. 201813 (7), 1521– 1538,  DOI: 10.1038/s41596-018-0003-zViewGoogle Scholar
  73. 73Young, T., III. An essay on the cohesion of fluids. Philosophical Transactions of the Royal Society of London 180595, 65– 87,  DOI: 10.1098/rstl.1805.0005ViewGoogle Scholar
  74. 74Kim, Y. C.; Kim, S.-H.; Kim, D.; Park, S.-J.; Park, J.-K. Plasma extraction in a capillary-driven microfluidic device using surfactant-added poly(dimethylsiloxane). Sens. Actuators, B 2010145 (2), 861– 868,  DOI: 10.1016/j.snb.2010.01.017ViewGoogle Scholar
  75. 75Washburn, E. W. The Dynamics of Capillary Flow. Physical Review 192117 (3), 273– 283,  DOI: 10.1103/PhysRev.17.273ViewGoogle Scholar
  76. 76Cito, S.; Ahn, Y. C.; Pallares, J.; Duarte, R. M.; Chen, Z.; Madou, M.; Katakis, I. Visualization and measurement of capillary-driven blood flow using spectral domain optical coherence tomography. Microfluid Nanofluidics 201213 (2), 227– 237,  DOI: 10.1007/s10404-012-0950-6ViewGoogle Scholar
  77. 77Berthier, E.; Dostie, A. M.; Lee, U. N.; Berthier, J.; Theberge, A. B. Open Microfluidic Capillary Systems. Anal Chem. 201991 (14), 8739– 8750,  DOI: 10.1021/acs.analchem.9b01429ViewGoogle Scholar
  78. 78Berthier, J.; Brakke, K. A.; Furlani, E. P.; Karampelas, I. H.; Poher, V.; Gosselin, D.; Cubizolles, M.; Pouteau, P. Whole blood spontaneous capillary flow in narrow V-groove microchannels. Sens. Actuators, B 2015206, 258– 267,  DOI: 10.1016/j.snb.2014.09.040ViewGoogle Scholar
  79. 79Hirt, C. W.; Nichols, B. D. Volume of fluid (VOF) method for the dynamics of free boundaries. J. Comput. Phys. 198139 (1), 201– 225,  DOI: 10.1016/0021-9991(81)90145-5ViewGoogle Scholar
  80. 80Chen, J.-L.; Shih, W.-H.; Hsieh, W.-H. AC electro-osmotic micromixer using a face-to-face, asymmetric pair of planar electrodes. Sens. Actuators, B 2013188, 11– 21,  DOI: 10.1016/j.snb.2013.07.012ViewGoogle Scholar
  81. 81Zhao, C.; Yang, C. Electrokinetics of non-Newtonian fluids: A review. Advances in Colloid and Interface Science 2013201-202, 94– 108,  DOI: 10.1016/j.cis.2013.09.001ViewGoogle Scholar
  82. 82Oh, K. W. 6 – Lab-on-chip (LOC) devices and microfluidics for biomedical applications. In MEMS for Biomedical Applications; Bhansali, S., Vasudev, A., Eds.; Woodhead Publishing, 2012; pp 150– 171.ViewGoogle Scholar
  83. 83Bello, M. S.; De Besi, P.; Rezzonico, R.; Righetti, P. G.; Casiraghi, E. Electroosmosis of polymer solutions in fused silica capillaries. ELECTROPHORESIS 199415 (1), 623– 626,  DOI: 10.1002/elps.1150150186ViewGoogle Scholar
  84. 84Park, H. M.; Lee, W. M. Effect of viscoelasticity on the flow pattern and the volumetric flow rate in electroosmotic flows through a microchannel. Lab Chip 20088 (7), 1163– 1170,  DOI: 10.1039/b800185eViewGoogle Scholar
  85. 85Afonso, A. M.; Alves, M. A.; Pinho, F. T. Analytical solution of mixed electro-osmotic/pressure driven flows of viscoelastic fluids in microchannels. J. Non-Newtonian Fluid Mech. 2009159 (1), 50– 63,  DOI: 10.1016/j.jnnfm.2009.01.006ViewGoogle Scholar
  86. 86Sousa, J. J.; Afonso, A. M.; Pinho, F. T.; Alves, M. A. Effect of the skimming layer on electro-osmotic─Poiseuille flows of viscoelastic fluids. Microfluid. Nanofluid. 201110 (1), 107– 122,  DOI: 10.1007/s10404-010-0651-yViewGoogle Scholar
  87. 87Zhao, C.; Yang, C. Electro-osmotic mobility of non-Newtonian fluids. Biomicrofluidics 20115 (1), 014110,  DOI: 10.1063/1.3571278ViewGoogle Scholar
  88. 88Pimenta, F.; Alves, M. A. Electro-elastic instabilities in cross-shaped microchannels. J. Non-Newtonian Fluid Mech. 2018259, 61– 77,  DOI: 10.1016/j.jnnfm.2018.04.004ViewGoogle Scholar
  89. 89Bezerra, W. S.; Castelo, A.; Afonso, A. M. Numerical Study of Electro-Osmotic Fluid Flow and Vortex Formation. Micromachines (Basel) 201910 (12), 796,  DOI: 10.3390/mi10120796ViewGoogle Scholar
  90. 90Ji, J.; Qian, S.; Liu, Z. Electroosmotic Flow of Viscoelastic Fluid through a Constriction Microchannel. Micromachines (Basel) 202112 (4), 417,  DOI: 10.3390/mi12040417ViewGoogle Scholar
  91. 91Zhao, C.; Yang, C. Exact solutions for electro-osmotic flow of viscoelastic fluids in rectangular micro-channels. Applied Mathematics and Computation 2009211 (2), 502– 509,  DOI: 10.1016/j.amc.2009.01.068ViewGoogle Scholar
  92. 92Gerum, R.; Mirzahossein, E.; Eroles, M.; Elsterer, J.; Mainka, A.; Bauer, A.; Sonntag, S.; Winterl, A.; Bartl, J.; Fischer, L. Viscoelastic properties of suspended cells measured with shear flow deformation cytometry. Elife 202211, e78823,  DOI: 10.7554/eLife.78823ViewGoogle Scholar
  93. 93Sadek, S. H.; Pinho, F. T.; Alves, M. A. Electro-elastic flow instabilities of viscoelastic fluids in contraction/expansion micro-geometries. J. Non-Newtonian Fluid Mech. 2020283, 104293,  DOI: 10.1016/j.jnnfm.2020.104293ViewGoogle Scholar
  94. 94Spanjaards, M.; Peters, G.; Hulsen, M.; Anderson, P. Numerical Study of the Effect of Thixotropy on Extrudate Swell. Polymers 202113 (24), 4383,  DOI: 10.3390/polym13244383ViewGoogle Scholar
  95. 95Rashidi, S.; Bafekr, H.; Valipour, M. S.; Esfahani, J. A. A review on the application, simulation, and experiment of the electrokinetic mixers. Chemical Engineering and Processing – Process Intensification 2018126, 108– 122,  DOI: 10.1016/j.cep.2018.02.021ViewGoogle Scholar
  96. 96Matsubara, K.; Narumi, T. Microfluidic mixing using unsteady electroosmotic vortices produced by a staggered array of electrodes. Chemical Engineering Journal 2016288, 638– 647,  DOI: 10.1016/j.cej.2015.12.013ViewGoogle Scholar
  97. 97Qaderi, A.; Jamaati, J.; Bahiraei, M. CFD simulation of combined electroosmotic-pressure driven micro-mixing in a microchannel equipped with triangular hurdle and zeta-potential heterogeneity. Chemical Engineering Science 2019199, 463– 477,  DOI: 10.1016/j.ces.2019.01.034ViewGoogle Scholar
  98. 98Cho, C.-C.; Chen, C.-L.; Chen, C. o.-K. Mixing enhancement in crisscross micromixer using aperiodic electrokinetic perturbing flows. International Journal of Heat and Mass Transfer 201255 (11), 2926– 2933,  DOI: 10.1016/j.ijheatmasstransfer.2012.02.006ViewGoogle Scholar
  99. 99Zhao, W.; Yang, F.; Wang, K.; Bai, J.; Wang, G. Rapid mixing by turbulent-like electrokinetic microflow. Chemical Engineering Science 2017165, 113– 121,  DOI: 10.1016/j.ces.2017.02.027ViewGoogle Scholar
  100. 100Tran, T.; Chakraborty, P.; Guttenberg, N.; Prescott, A.; Kellay, H.; Goldburg, W.; Goldenfeld, N.; Gioia, G. Macroscopic effects of the spectral structure in turbulent flows. Nat. Phys. 20106 (6), 438– 441,  DOI: 10.1038/nphys1674ViewGoogle Scholar
  101. 101Toner, M.; Irimia, D. Blood-on-a-chip. Annu. Rev. Biomed Eng. 20057, 77– 103,  DOI: 10.1146/annurev.bioeng.7.011205.135108ViewGoogle Scholar
  102. 102Maria, M. S.; Rakesh, P. E.; Chandra, T. S.; Sen, A. K. Capillary flow of blood in a microchannel with differential wetting for blood plasma separation and on-chip glucose detection. Biomicrofluidics 201610 (5), 054108,  DOI: 10.1063/1.4962874ViewGoogle Scholar
  103. 103Tripathi, S.; Varun Kumar, Y. V. B.; Prabhakar, A.; Joshi, S. S.; Agrawal, A. Passive blood plasma separation at the microscale: a review of design principles and microdevices. Journal of Micromechanics and Microengineering 201525 (8), 083001,  DOI: 10.1088/0960-1317/25/8/083001ViewGoogle Scholar
  104. 104Mohammadi, M.; Madadi, H.; Casals-Terré, J. Microfluidic point-of-care blood panel based on a novel technique: Reversible electroosmotic flow. Biomicrofluidics 20159 (5), 054106,  DOI: 10.1063/1.4930865ViewGoogle Scholar
  105. 105Kang, D. H.; Kim, K.; Kim, Y. J. An anti-clogging method for improving the performance and lifespan of blood plasma separation devices in real-time and continuous microfluidic systems. Sci. Rep 20188 (1), 17015,  DOI: 10.1038/s41598-018-35235-4ViewGoogle Scholar
  106. 106Li, Z.; Pollack, G. H. Surface-induced flow: A natural microscopic engine using infrared energy as fuel. Science Advances 20206 (19), eaba0941  DOI: 10.1126/sciadv.aba0941ViewGoogle Scholar
  107. 107Mercado-Uribe, H.; Guevara-Pantoja, F. J.; García-Muñoz, W.; García-Maldonado, J. S.; Méndez-Alcaraz, J. M.; Ruiz-Suárez, J. C. On the evolution of the exclusion zone produced by hydrophilic surfaces: A contracted description. J. Chem. Phys. 2021154 (19), 194902,  DOI: 10.1063/5.0043084ViewGoogle Scholar
  108. 108Yalcin, O.; Jani, V. P.; Johnson, P. C.; Cabrales, P. Implications Enzymatic Degradation of the Endothelial Glycocalyx on the Microvascular Hemodynamics and the Arteriolar Red Cell Free Layer of the Rat Cremaster Muscle. Front Physiol 20189, 168,  DOI: 10.3389/fphys.2018.00168ViewGoogle Scholar

FLOW DEM

FLOW-3D DEM Module 개요

FLOW DEM 은 FLOW-3D 의 기체 및 액체 유동 해석에 DEM(Discrete Element Method : 개별 요소법)공법인 입자의 거동을 분석해주는 모듈입니다.

dem9

dem10
주요 기능 :고체 요소의 충돌, 스프링(Spring) / 대시 포트(Dash Pot) 모델 적용 Void, 1 fluid, 2 fluid(자유 계면 포함) 각각의 모드에 대응 가변 밀도 / 가변 직경 입자 크기조절로 입자 특성을 유지하면서 입자 수를 감소 독립적인 DEM의 Sub Time Step 이용

Discrete Element Method : 개별 요소법

다수의 고체 요소의 충돌 운동을 분석하는 데 유용합니다. 유동 해석과 함께 사용하면 광범위한 용도에 응용을 할 수 있습니다.

dem1

입자 간의 충돌

Voigt model은 스프링(Spring) 및 대시 포트(Dash pot)의 조합에 의해 입자 충돌 시의 힘을 평가합니다. 탄성력 부분은 스프링 모델에서,
비탄성 충돌의 에너지 소산부분은 대시 포트 모델에서 시뮬레이션되고 있으며, 중량 및 항력은 작용하는 외력으로 고려 될 수 있습니다.

분석 모드

기본적으로 이용하는 운동 방정식은 FLOW-3D 에 사용되는 질량 입자의 운동 방정식과 같은 것이지만, 여기에 DEM으로
평가되는 항목이 추가되기 형태로되어 있으며, 실제 시뮬레이션으로는 ‘void + DEM’, ‘1 Fluid + DEM’ , ‘ 1 Fluid 자유계면 + DEM ‘을 기본 유동 모드로 취급이 가능합니다.

dem4

입자 유형

입자 타입도 표준 기능의 질량 입자 모델처럼 입자 크기 (반경)와 밀도가 동일한 것 외, 크기는 같지만 밀도가 다른 것이나 밀도는 같지만 크기가 다른 것 등도 취급 가능합니다. 이로 인해 표준 질량 입자 모델에서는 입자 간의 상호 작용이 고려되어 있지 않기 때문에 모든 아래에 가라 앉아 버리고 있었지만, FLOW DEM을 이용하여 기하학적 관계를 평가하는 것이 가능합니다.

dem7

응용 분야

1. Mechanical Engineering 분야

수지 충전, 스쿠류 이송, 분말 이송 / Resin filling, screw conveyance, powder conveyance

2. Civil Engineering분야

3. Civil Engineering 분야

파편, 자갈, 낙 성/ Debris flow, gravel, falling rock

dem11

3. Chemical Engineering, Pharmaceutics 분야

유동층, 사이클론, 교반기 / Fluidized bed, cyclone, stirrer

dem12

4. MEMS, Electrical Engineering 분야

하전 입자를 포함한 전기장 해석 등

dem15

입자 그룹 가시화

그룹 가시화

DEM은 일반적으로 다수의 입자를 필요로하는 분석을 상정하고 있습니다. 
다만 이 경우, 계산 부하가 높아 지므로 현실적인 계산자원을 고려하면, 입자 수가 너무 많아 현실적으로 취급 할 수 없는 경우 입자의 특성은 유지하고 숫자를 줄여 가시화할 필요가 있습니다 .
일반적인 유동해석 계산의 메쉬 해상도에 해당합니다.
메쉬 수 많음 (계산 부하 큼) → 소 (계산 부하 적음)
입자 수 다 (계산 부하 큼) → 소 (계산 부하 적음)

원래 입자수

입자 사이즈를 키운경우

그룹 가시화

  • 입자 수를 줄이기 위해 그대로 입경을 크게했을 경우와 그룹 가시화 한 경우의 비교.
  • 입자 크기를 크게하면 개별 입자 특성이 달라지기 때문에 거동이 달라진다. (본 사례에서는 부력이 커진다.)
  • 그룹 가시화의 경우 개별 특성은 동일 원래의 거동과 대체로 일치한다.

주조 시뮬레이션에 DEM 적용

그룹 가시화 비교 예

그룹 가시화한 경우와 입경을 크게하여 수를 줄인 경우, 입경을 크게하면
개별 입자 특성이 변화하여 거동이 바뀌어 버리기 때문에 실제 계산으로는 사용할 수 어렵습니다.

중자 모래 분사 분석

DEM에서의 계산부하를 생각할 때는 입자모델에 의한 안정제한을 고려해야 하지만 서브타임스텝이라는 개념을 도입함으로써 입자의 경우와 유체의 경우의 타임스텝을 바꾸고 필요이상으로 계산시간을 들이지 않고 효율적으로 계산하는 것을 가능하게 하고 있습니다.

이를 통해 예를 들어 중자사 분사 시뮬레이션 실험에서는 이러한 문제로 자주 이용되는 빙엄 유체에서는 실험과의 정합성이 별로 좋지 않기 때문에 당사에서는 이전부터 입상류 모델이라는 모델을 개발하고 연속체로부터의 접근에서도 실험과의 높은 정합성을 실현할 수 있는 모델화를 해왔는데, 이번에 DEM을 사용해도 그것과 거의 같은 결과를 얻습니다. 할 수 있음을 확인할 수 있었다.

Reference :

  • Lefebvre D., Mackenbrock A., Vidal V., Pavan V. and Haigh PM, 2004,
  • Development and use of simulation in the Design of Blown Cores and Moulds

FLOW-3D AM

flow3d AM-product
FLOW-3D AM-product

와이어 파우더 기반 DED | Wire Powder Based DED

일부 연구자들은 부품을 만들기 위해 더 넓은 범위의 처리 조건을 사용하여 하이브리드 와이어 분말 기반 DED 시스템을 찾고 있습니다. 예를 들어, 이 시뮬레이션은 다양한 분말 및 와이어 이송 속도를 가진 하이브리드 시스템을 살펴봅니다.

와이어 기반 DED | Wire Based DED

와이어 기반 DED는 분말 기반 DED보다 처리량이 높고 낭비가 적지만 재료 구성 및 증착 방향 측면에서 유연성이 떨어집니다. FLOW-3D AM 은 와이어 기반 DED의 처리 결과를 이해하는데 유용하며 최적화 연구를 통해 빌드에 대한 와이어 이송 속도 및 직경과 같은 최상의 처리 매개 변수를 찾을 수 있습니다.

FLOW-3D AM은 레이저 파우더 베드 융합 (L-PBF), 바인더 제트 및 DED (Directed Energy Deposition)와 같은 적층 제조 공정 ( additive manufacturing )을 시뮬레이션하고 분석하는 CFD 소프트웨어입니다. FLOW-3D AM 의 다중 물리 기능은 공정 매개 변수의 분석 및 최적화를 위해 분말 확산 및 압축, 용융 풀 역학, L-PBF 및 DED에 대한 다공성 형성, 바인더 분사 공정을 위한 수지 침투 및 확산에 대해 매우 정확한 시뮬레이션을 제공합니다.

3D 프린팅이라고도하는 적층 제조(additive manufacturing)는 일반적으로 층별 접근 방식을 사용하여, 분말 또는 와이어로 부품을 제조하는 방법입니다. 금속 기반 적층 제조 공정에 대한 관심은 지난 몇 년 동안 시작되었습니다. 오늘날 사용되는 3 대 금속 적층 제조 공정은 PBF (Powder Bed Fusion), DED (Directed Energy Deposition) 및 바인더 제트 ( Binder jetting ) 공정입니다.  FLOW-3D  AM  은 이러한 각 프로세스에 대한 고유 한 시뮬레이션 통찰력을 제공합니다.

파우더 베드 융합 및 직접 에너지 증착 공정에서 레이저 또는 전자 빔을 열원으로 사용할 수 있습니다. 두 경우 모두 PBF용 분말 형태와 DED 공정용 분말 또는 와이어 형태의 금속을 완전히 녹여 융합하여 층별로 부품을 형성합니다. 그러나 바인더 젯팅(Binder jetting)에서는 결합제 역할을 하는 수지가 금속 분말에 선택적으로 증착되어 층별로 부품을 형성합니다. 이러한 부품은 더 나은 치밀화를 달성하기 위해 소결됩니다.

FLOW-3D AM 의 자유 표면 추적 알고리즘과 다중 물리 모델은 이러한 각 프로세스를 높은 정확도로 시뮬레이션 할 수 있습니다. 레이저 파우더 베드 융합 (L-PBF) 공정 모델링 단계는 여기에서 자세히 설명합니다. DED 및 바인더 분사 공정에 대한 몇 가지 개념 증명 시뮬레이션도 표시됩니다.

레이저 파우더 베드 퓨전 (L-PBF)

LPBF 공정에는 유체 흐름, 열 전달, 표면 장력, 상 변화 및 응고와 같은 복잡한 다중 물리학 현상이 포함되어 공정 및 궁극적으로 빌드 품질에 상당한 영향을 미칩니다. FLOW-3D AM 의 물리적 모델은 질량, 운동량 및 에너지 보존 방정식을 동시에 해결하는 동시에 입자 크기 분포 및 패킹 비율을 고려하여 중규모에서 용융 풀 현상을 시뮬레이션합니다.

FLOW-3D DEM FLOW-3D WELD 는 전체 파우더 베드 융합 공정을 시뮬레이션하는 데 사용됩니다. L-PBF 공정의 다양한 단계는 분말 베드 놓기, 분말 용융 및 응고,이어서 이전에 응고 된 층에 신선한 분말을 놓는 것, 그리고 다시 한번 새 층을 이전 층에 녹이고 융합시키는 것입니다. FLOW-3D AM  은 이러한 각 단계를 시뮬레이션하는 데 사용할 수 있습니다.

파우더 베드 부설 공정

FLOW-3D DEM을 통해 분말 크기 분포, 재료 특성, 응집 효과는 물론 롤러 또는 블레이드 움직임 및 상호 작용과 같은 기하학적 효과와 관련된 분말 확산 및 압축을 이해할 수 있습니다. 이러한 시뮬레이션은 공정 매개 변수가 후속 인쇄 공정에서 용융 풀 역학에 직접적인 영향을 미치는 패킹 밀도와 같은 분말 베드 특성에 어떻게 영향을 미치는지에 대한 정확한 이해를 제공합니다.

다양한 파우더 베드 압축을 달성하는 한 가지 방법은 베드를 놓는 동안 다양한 입자 크기 분포를 선택하는 것입니다. 아래에서 볼 수 있듯이 세 가지 크기의 입자 크기 분포가 있으며, 이는 가장 높은 압축을 제공하는 Case 2와 함께 다양한 분말 베드 압축을 초래합니다.

파우더 베드 분포 다양한 입자 크기 분포
세 가지 다른 입자 크기 분포를 사용하여 파우더 베드 배치
파우더 베드 압축 결과
세 가지 다른 입자 크기 분포를 사용한 분말 베드 압축

입자-입자 상호 작용, 유체-입자 결합 및 입자 이동 물체 상호 작용은 FLOW-3D DEM을 사용하여 자세히 분석 할 수도 있습니다 . 또한 입자간 힘을 지정하여 분말 살포 응용 분야를 보다 정확하게 연구 할 수도 있습니다.

FLOW-3D AM  시뮬레이션은 이산 요소 방법 (DEM)을 사용하여 역 회전하는 원통형 롤러로 인한 분말 확산을 연구합니다. 비디오 시작 부분에서 빌드 플랫폼이 위로 이동하는 동안 분말 저장소가 아래로 이동합니다. 그 직후, 롤러는 분말 입자 (초기 위치에 따라 색상이 지정됨)를 다음 층이 녹고 구축 될 준비를 위해 구축 플랫폼으로 펼칩니다. 이러한 시뮬레이션은 저장소에서 빌드 플랫폼으로 전송되는 분말 입자의 선호 크기에 대한 추가 통찰력을 제공 할 수 있습니다.

Melting | 파우더 베드 용해

DEM 시뮬레이션에서 파우더 베드가 생성되면 STL 파일로 추출됩니다. 다음 단계는 CFD를 사용하여 레이저 용융 공정을 시뮬레이션하는 것입니다. 여기서는 레이저 빔과 파우더 베드의 상호 작용을 모델링 합니다. 이 프로세스를 정확하게 포착하기 위해 물리학에는 점성 흐름, 용융 풀 내의 레이저 반사 (광선 추적을 통해), 열 전달, 응고, 상 변화 및 기화, 반동 압력, 차폐 가스 압력 및 표면 장력이 포함됩니다. 이 모든 물리학은 이 복잡한 프로세스를 정확하게 시뮬레이션하기 위해 TruVOF 방법을 기반으로 개발되었습니다.

레이저 출력 200W, 스캔 속도 3.0m / s, 스폿 반경 100μm에서 파우더 베드의 용융 풀 분석.

용융 풀이 응고되면 FLOW-3D AM  압력 및 온도 데이터를 Abaqus 또는 MSC Nastran과 같은 FEA 도구로 가져와 응력 윤곽 및 변위 프로파일을 분석 할 수도 있습니다.

Multilayer | 다층 적층 제조

용융 풀 트랙이 응고되면 DEM을 사용하여 이전에 응고된 층에 새로운 분말 층의 확산을 시뮬레이션 할 수 있습니다. 유사하게, 레이저 용융은 새로운 분말 층에서 수행되어 후속 층 간의 융합 조건을 분석 할 수 있습니다.

해석 진행 절차는 첫 번째 용융층이 응고되면 입자의 두 번째 층이 응고 층에 증착됩니다. 새로운 분말 입자 층에 레이저 공정 매개 변수를 지정하여 용융 풀 시뮬레이션을 다시 수행합니다. 이 프로세스를 여러 번 반복하여 연속적으로 응고된 층 간의 융합, 빌드 내 온도 구배를 평가하는 동시에 다공성 또는 기타 결함의 형성을 모니터링 할 수 있습니다.

다층 적층 적층 제조 시뮬레이션

LPBF의 키홀 링 | Keyholing in LPBF

키홀링 중 다공성은 어떻게 형성됩니까? 이것은 TU Denmark의 연구원들이 FLOW-3D AM을 사용하여 답변한 질문이었습니다. 레이저 빔의 적용으로 기판이 녹으면 기화 및 상 변화로 인한 반동 압력이 용융 풀을 압박합니다. 반동 압력으로 인한 하향 흐름과 레이저 반사로 인한 추가 레이저 에너지 흡수가 공존하면 폭주 효과가 발생하여 용융 풀이 Keyholing으로 전환됩니다. 결국, 키홀 벽을 따라 온도가 변하기 때문에 표면 장력으로 인해 벽이 뭉쳐져서 진행되는 응고 전선에 의해 갇힐 수 있는 공극이 생겨 다공성이 발생합니다. FLOW-3D AM 레이저 파우더 베드 융합 공정 모듈은 키홀링 및 다공성 형성을 시뮬레이션 하는데 필요한 모든 물리 모델을 보유하고 있습니다.

바인더 분사 (Binder jetting)

Binder jetting 시뮬레이션은 모세관 힘의 영향을받는 파우더 베드에서 바인더의 확산 및 침투에 대한 통찰력을 제공합니다. 공정 매개 변수와 재료 특성은 증착 및 확산 공정에 직접적인 영향을 미칩니다.

Scan Strategy | 스캔 전략

스캔 전략은 온도 구배 및 냉각 속도에 영향을 미치기 때문에 미세 구조에 직접적인 영향을 미칩니다. 연구원들은 FLOW-3D AM 을 사용하여 결함 형성과 응고된 금속의 미세 구조에 영향을 줄 수 있는 트랙 사이에서 발생하는 재 용융을 이해하기 위한 최적의 스캔 전략을 탐색하고 있습니다. FLOW-3D AM 은 하나 또는 여러 레이저에 대해 시간에 따른 방향 속도를 구현할 때 완전한 유연성을 제공합니다.

Beam Shaping | 빔 형성

레이저 출력 및 스캔 전략 외에도 레이저 빔 모양과 열유속 분포는 LPBF 공정에서 용융 풀 역학에 큰 영향을 미칩니다. AM 기계 제조업체는 공정 안정성 및 처리량에 대해 다중 코어 및 임의 모양의 레이저 빔 사용을 모색하고 있습니다. FLOW-3D AM을 사용하면 멀티 코어 및 임의 모양의 빔 프로파일을 구현할 수 있으므로 생산량을 늘리고 부품 품질을 개선하기 위한 최상의 구성에 대한 통찰력을 제공 할 수 있습니다.

이 영역에서 수행 된 일부 작업에 대해 자세히 알아 보려면 “The Next Frontier of Metal AM”웨비나를 시청하십시오.

Multi-material Powder Bed Fusion | 다중 재료 분말 베드 융합

이 시뮬레이션에서 스테인리스 강 및 알루미늄 분말은 FLOW-3D AM 이 용융 풀 역학을 정확하게 포착하기 위해 추적하는 독립적으로 정의 된 온도 의존 재료 특성을 가지고 있습니다. 시뮬레이션은 용융 풀에서 재료 혼합을 이해하는 데 도움이됩니다.

다중 재료 용접 사례 연구

이종 금속의 레이저 키홀 용접에서 금속 혼합 조사

GM과 University of Utah의 연구원들은 FLOW-3D WELD 를 사용 하여 레이저 키홀 용접을 통한 이종 금속의 혼합을 이해했습니다. 그들은 반동 압력 및 Marangoni 대류와 관련하여 구리와 알루미늄의 혼합 농도에 대한 레이저 출력 및 스캔 속도의 영향을 조사했습니다. 그들은 시뮬레이션을 실험 결과와 비교했으며 샘플 내의 절단 단면에서 재료 농도 사이에 좋은 일치를 발견했습니다.

이종 금속의 레이저 키홀 용접에서 금속 혼합 조사
이종 금속의 레이저 키홀 용접에서 금속 혼합 조사
참조 : Wenkang Huang, Hongliang Wang, Teresa Rinker, Wenda Tan, 이종 금속의 레이저 키홀 용접에서 금속 혼합 조사 , Materials & Design, Volume 195, (2020). https://doi.org/10.1016/j.matdes.2020.109056
참조 : Wenkang Huang, Hongliang Wang, Teresa Rinker, Wenda Tan, 이종 금속의 레이저 키홀 용접에서 금속 혼합 조사 , Materials & Design, Volume 195, (2020). https://doi.org/10.1016/j.matdes.2020.109056

방향성 에너지 증착

FLOW-3D AM 의 내장 입자 모델 을 사용하여 직접 에너지 증착 프로세스를 시뮬레이션 할 수 있습니다. 분말 주입 속도와 고체 기질에 입사되는 열유속을 지정함으로써 고체 입자는 용융 풀에 질량, 운동량 및 에너지를 추가 할 수 있습니다. 다음 비디오에서 고체 금속 입자가 용융 풀에 주입되고 기판에서 용융 풀의 후속 응고가 관찰됩니다.

Fig.1 Schematic diagram of the novel cytometric device

Fabrication and Experimental Investigation of a Novel 3D Hydrodynamic Focusing Micro Cytometric Device

Yongquan Wang*a , Jingyuan Wangb, Hualing Chenc

School of Mechanical Engineering, Xi’an Jiaotong University, Xi’an, Shaanxi, 710049, P. R. China
a yqwang@mail.xjtu.edu.cn,, bwjy2006@stu.xjtu.edu.cn,, c hlchen@mail.xjtu.edu.cn,

Abstract:

This paper presents the fabrication of a novel micro-machined cytometric device, and the experimental investigations for its 3D hydrodynamic focusing performance. The proposed device is simple in structure, with the uniqueness that the depth of its microchannels is non-uniform. Using the SU-8 soft lithography containing two exposures, as well as micro-molding techniques, the PDMS device is successfully fabricated. Two kinds of experiments, i.e., the red ink fluidity observation experiments and the fluorescent optical experiments, are then performed for the device prototypes with different step heights, or channel depth differences, to explore the influence laws of the feature parameter on the devices hydrodynamic focusing behaviors. The experimental results show that the introducing of the steps can efficiently enhance the vertical focusing performance of the device. At appropriate geometry and operating conditions, good 3D hydrodynamic focusing can be obtained.

Korea Abstract

이 논문은 새로운 마이크로 머신 세포 측정 장치의 제조와 3D 유체 역학적 초점 성능에 대한 실험적 조사를 제시합니다. 제안 된 장치는 구조가 단순하며, 마이크로 채널의 깊이가 균일하지 않다는 독특함이 있습니다. 두 가지 노출이 포함 된 SU-8 소프트 리소그래피와 마이크로 몰딩 기술을 사용하여 PDMS 장치가 성공적으로 제작되었습니다. 그런 다음 두 종류의 실험, 즉 적색 잉크 유동성 관찰 실험과 형광 광학 실험을 단계 높이 또는 채널 깊이 차이가 다른 장치 프로토 타입에 대해 수행하여 장치 유체 역학적 초점에 대한 기능 매개 변수의 영향 법칙을 탐색합니다. 행동. 실험 결과는 단계의 도입이 장치의 수직 초점 성능을 효율적으로 향상시킬 수 있음을 보여줍니다. 적절한 형상과 작동 조건에서 우수한 3D 유체 역학적 초점을 얻을 수 있습니다.

Keywords

Flow cytometer, Hydrodynamic focusing, Three-dimensional (3D), Micro-machined

Fig.1 Schematic diagram of the novel cytometric device
Fig.1 Schematic diagram of the novel cytometric device
Fig.2 Overview of the cytometric device fabrication process
Fig.2 Overview of the cytometric device fabrication process
Fig.3 The fabricated micro cytometric device Fig.4 Experiment setup for focusing performance
Fig.3 The fabricated micro cytometric device Fig. 4 Experiment setup for focusing performance
Fig.5 Horizontal focusing images of two devices with and without steps
Fig.5 Horizontal focusing images of two devices with and without steps
Fig.6 Channel cross-section fluorescence images for different step heights
Fig.6 Channel cross-section fluorescence images for different step heights

References 

Fig.7 Effect of the step height on the 3D focusing at different velocity ratios
Fig.7 Effect of the step height on the 3D focusing at different velocity ratios

Conclusions

In this paper, we presented a novel micro-machined cytometric device and its fabrication process,
emphasizing on the experimental investigations for its 3D hydrodynamic focusing performance. The
proposed device is simple in structure, low cost, and easy to be batch produced. Besides this, as a
device based on standard micro-fabrication methodology, it can be conveniently integrated with other
micro-fluidic and/or micro-optical units to form a complete detection and analysis system.
The experimental tests for the prototype devices not only verified the design conception, but also
gave us a comprehensive understanding of the device hydro-focusing performance. The experimental
results show that, as the uniqueness of this design, the introducing of the feature steps can
significantly enhance the vertical focusing performance of the devices, which is crucial for the
achievement of 3D focusing. In summary, for the proposed novel device, good 3D hydrodynamic
focusing can be attained at appropriate geometry and operating conditions.
In addition, an improved design can be obtained by replacing the flat cover with an identical
device unit, in other words, the same two device units are bonded together (The channels are inward
and aligned) to form a new device. Then the sample stream can focused to the center of the assembly
outlet channel due to the hydrodynamic forces equally in both horizontal and vertical directions, and
thus avoiding the adsorption or friction issues of cells/particles to the top channel wall.

References

[1] Mandy FF, Bergeron M, Minkus T, Principles of flow cytometry. Transfusion Science Transfusion Science, 16 (1995) 303.

DOI: 10.1016/0955-3886(95)00041-0

[2] Rieseberg M, Kasper C, Reardon KF, and Scheper T, Flow cytometry in biotechnology, Appl. Microbiol. Biotechnol. 56 (2001) 350.

[3] Chung TD, Kim HC, Recent advances in miniaturized microfluidic flow cytometry for clinical use. Electrophoresis, Electrophoresis. 28(2007) 4511.

DOI: 10.1002/elps.200700620

[4] Xuan X, Zhu J, Church C, Particle focusing in microfluidic devices, Microfluid Nanofluid. 9(2010) 1-16.

DOI: 10.1007/s10404-010-0602-7

[5] Stone H A, Stroock A D and Ajdari A, Engineering flows in small devices: Microfluidics toward a lab-on-a-chip, Annu. Rev. Fluid Mech. 36 (2004) 381-411.

[6] Fu LM, Yang RJ, Lin CH, Pan YJ, and Lee GB, Electrokinetically-driven microflow cytometers with integrated fiber optics for on-line cell/particle collection, Analytica Chimica Acta. 507(2004) 163-169.

DOI: 10.1016/j.aca.2003.10.028

[7] Applegate Jr RW, Schafer DN, Amir W, Squier J, Vestad T, Oakey J and Marr DWM, Optically integrated microfluidic systems for cellular characterization and manipulation, J. Opt. A: Pure Appl. Opt. 9(2007) 122-128.

DOI: 10.1088/1464-4258/9/8/s03

[8] Chang CM, Hsiung SK, Lee GB, Micro flow cytometer chip integrated with micro-pumps/micro-valves for multi-wavelength cell counting and sorting, Jpn. J. Appl. Phys. 46 (2007): 3126-3134.

DOI: 10.1143/jjap.46.3126

[9] Lee GB, Hung CI, Ke BJ, Huang GR, Hwei BH, and Lai Hui-Fang, Hydrodynamic focusing for a micromachined flow cytometer, J Fluids Engineering 123(2001) 672-679.

DOI: 10.1115/1.1385514

[10] Weigl BH, Bardell R, Schulte T, Battrell F and Hayenga J, Design and rapid prototyping of thin-film laminate-based microfluidic devices, Biomed Microdevices. 3(2001) 267-274.

DOI: 10.1023/a:1012448412811

[11] Yang AS, Hsieh WH, Hydrodynamic focusing investigation in a micro-flow cytometer, Biomed Microdevices, 9(2007) 113-122.

DOI: 10.1007/s10544-006-9003-9

[12] Ateya DA, Erickson JS, Howell PB Jr, Hilliard LR, Golden JP, Ligler FS, The good, the bad, and the tiny: a review of microflow cytometry, Anal Bioanal Chem. 391(2008) 1485-1498.

DOI: 10.1007/s00216-007-1827-5

[13] Goranovic G, Perch-Nielsen I, Larsen UD, Wolff A, Kutter J and Telleman P, Three-Dimensional Single Step Flow Sheathing in Micro Cell Sorters, Proceedings of MSM Conference. (2001) pp.242-245.

[14] Lin CH, Lee GB, Fu LM, and Hwey BH, Vertical focusing device utilizing dielectrophoretic force and its application on mocroflow cytometer, J. Microelectromech. Syst. 13 (2004) 923-932.

DOI: 10.1109/jmems.2004.838352

[15] Yang R, Feeback DL, Wang W, Microfabrication and test of a three-dimensional polymer hydro-focusing unit for flow cytometry applications, Sens. Actuat. A. 118(2005) 259-267.

DOI: 10.1016/j.sna.2004.09.001

[16] Hairer G, Pärr GS, Svasek P, Jachimowicz A, and Vellekoop MJ, Investigations of micrometer sample stream profiles in a three-dimensional hydrodynamic focusing device, Sens. Actuat. B. 132 (2008) 518-524.

DOI: 10.1016/j.snb.2007.11.018

[17] Mao X, Lin SC, Dong C, and Huang TJ, Single-layer planar on-chip flow cytometer using microfluidic drifting based three-dimensional (3D) hydrodynamic focusing, Lab Chip. 9 (2009) 1583-1589.

DOI: 10.1039/b820138b

[18] Wang Y, Wang J, Chen H, Zhu Z, and Wang B, Prototype of a novel micro-machined cytometer and its 3D hydrodynamic focusing properties, Microsyst. Technol. 18(2012) 1991-(1997).

DOI: 10.1007/s00542-012-1525-x

Figure 1. (a) Top view of the microfluidic-magnetophoretic device, (b) Schematic representation of the channel cross-sections studied in this work, and (c) the magnet position relative to the channel location (Sepy and Sepz are the magnet separation distances in y and z, respectively).

Continuous-Flow Separation of Magnetic Particles from Biofluids: How Does the Microdevice Geometry Determine the Separation Performance?

1Department of Chemical and Biomolecular Engineering, ETSIIT, University of Cantabria, Avda. Los Castros s/n, 39005 Santander, Spain
2William G. Lowrie Department of Chemical and Biomolecular Engineering, The Ohio State University, 151 W. Woodruff Ave., Columbus, OH 43210, USA
*Author to whom correspondence should be addressed.
Sensors 202020(11), 3030; https://doi.org/10.3390/s20113030
Received: 16 April 2020 / Revised: 21 May 2020 / Accepted: 25 May 2020 / Published: 27 May 2020
(This article belongs to the Special Issue Lab-on-a-Chip and Microfluidic Sensors)

Abstract

The use of functionalized magnetic particles for the detection or separation of multiple chemicals and biomolecules from biofluids continues to attract significant attention. After their incubation with the targeted substances, the beads can be magnetically recovered to perform analysis or diagnostic tests. Particle recovery with permanent magnets in continuous-flow microdevices has gathered great attention in the last decade due to the multiple advantages of microfluidics. As such, great efforts have been made to determine the magnetic and fluidic conditions for achieving complete particle capture; however, less attention has been paid to the effect of the channel geometry on the system performance, although it is key for designing systems that simultaneously provide high particle recovery and flow rates. Herein, we address the optimization of Y-Y-shaped microchannels, where magnetic beads are separated from blood and collected into a buffer stream by applying an external magnetic field. The influence of several geometrical features (namely cross section shape, thickness, length, and volume) on both bead recovery and system throughput is studied. For that purpose, we employ an experimentally validated Computational Fluid Dynamics (CFD) numerical model that considers the dominant forces acting on the beads during separation. Our results indicate that rectangular, long devices display the best performance as they deliver high particle recovery and high throughput. Thus, this methodology could be applied to the rational design of lab-on-a-chip devices for any magnetically driven purification, enrichment or isolation.

Keywords: particle magnetophoresisCFDcross sectionchip fabrication

Korea Abstract

생체 유체에서 여러 화학 물질과 생체 분자의 검출 또는 분리를위한 기능화 된 자성 입자의 사용은 계속해서 상당한 관심을 받고 있습니다. 표적 물질과 함께 배양 한 후 비드를 자기 적으로 회수하여 분석 또는 진단 테스트를 수행 할 수 있습니다. 연속 흐름 마이크로 장치에서 영구 자석을 사용한 입자 회수는 마이크로 유체의 여러 장점으로 인해 지난 10 년 동안 큰 관심을 모았습니다. 

따라서 완전한 입자 포획을 달성하기 위한 자기 및 유체 조건을 결정하기 위해 많은 노력을 기울였습니다. 그러나 높은 입자 회수율과 유속을 동시에 제공하는 시스템을 설계하는 데있어 핵심이기는 하지만 시스템 성능에 대한 채널 형상의 영향에 대해서는 덜주의를 기울였습니다. 

여기에서 우리는 자기 비드가 혈액에서 분리되고 외부 자기장을 적용하여 버퍼 스트림으로 수집되는 YY 모양의 마이크로 채널의 최적화를 다룹니다. 비드 회수 및 시스템 처리량에 대한 여러 기하학적 특징 (즉, 단면 형상, 두께, 길이 및 부피)의 영향을 연구합니다. 

이를 위해 분리 중에 비드에 작용하는 지배적인 힘을 고려하는 실험적으로 검증 된 CFD (Computational Fluid Dynamics) 수치 모델을 사용합니다. 우리의 결과는 직사각형의 긴 장치가 높은 입자 회수율과 높은 처리량을 제공하기 때문에 최고의 성능을 보여줍니다. 

따라서 이 방법론은 자기 구동 정제, 농축 또는 분리를 위한 랩온어 칩 장치의 합리적인 설계에 적용될 수 있습니다.

Figure 1. (a) Top view of the microfluidic-magnetophoretic device, (b) Schematic representation of the channel cross-sections studied in this work, and (c) the magnet position relative to the channel location (Sepy and Sepz are the magnet separation distances in y and z, respectively).
Figure 1. (a) Top view of the microfluidic-magnetophoretic device, (b) Schematic representation of the channel cross-sections studied in this work, and (c) the magnet position relative to the channel location (Sepy and Sepz are the magnet separation distances in y and z, respectively).
Figure 2. (a) Channel-magnet configuration and (b–d) magnetic force distribution in the channel midplane for 2 mm, 5 mm and 10 mm long rectangular (left) and U-shaped (right) devices.
Figure 2. (a) Channel-magnet configuration and (b–d) magnetic force distribution in the channel midplane for 2 mm, 5 mm and 10 mm long rectangular (left) and U-shaped (right) devices.
Figure 3. (a) Velocity distribution in a section perpendicular to the flow for rectangular (left) and U-shaped (right) cross section channels, and (b) particle location in these cross sections.
Figure 3. (a) Velocity distribution in a section perpendicular to the flow for rectangular (left) and U-shaped (right) cross section channels, and (b) particle location in these cross sections.
Figure 4. Influence of fluid flow rate on particle recovery when the applied magnetic force is (a) different and (b) equal in U-shaped and rectangular cross section microdevices.
Figure 4. Influence of fluid flow rate on particle recovery when the applied magnetic force is (a) different and (b) equal in U-shaped and rectangular cross section microdevices.
Figure 5. Magnetic bead capture as a function of fluid flow rate for all of the studied geometries.
Figure 5. Magnetic bead capture as a function of fluid flow rate for all of the studied geometries.
Figure 6. Influence of (a) magnetic and fluidic forces (J parameter) and (b) channel geometry (θ parameter) on particle recovery. Note that U-2mm does not accurately fit a line.
Figure 6. Influence of (a) magnetic and fluidic forces (J parameter) and (b) channel geometry (θ parameter) on particle recovery. Note that U-2mm does not accurately fit a line.
Figure 7. Dependence of bead capture on the (a) functional channel volume and (b) particle residence time (tres). Note that in the curve fitting expressions V represents the functional channel volume and that U-2mm does not accurately fit a line.
Figure 7. Dependence of bead capture on the (a) functional channel volume and (b) particle residence time (tres). Note that in the curve fitting expressions V represents the functional channel volume and that U-2mm does not accurately fit a line.

References

  1. Gómez-Pastora, J.; Xue, X.; Karampelas, I.H.; Bringas, E.; Furlani, E.P.; Ortiz, I. Analysis of separators for magnetic beads recovery: From large systems to multifunctional microdevices. Sep. Purif. Technol. 2017172, 16–31. [Google Scholar] [CrossRef]
  2. Wise, N.; Grob, T.; Morten, K.; Thompson, I.; Sheard, S. Magnetophoretic velocities of superparamagnetic particles, agglomerates and complexes. J. Magn. Magn. Mater. 2015384, 328–334. [Google Scholar] [CrossRef]
  3. Khashan, S.A.; Elnajjar, E.; Haik, Y. CFD simulation of the magnetophoretic separation in a microchannel. J. Magn. Magn. Mater. 2011323, 2960–2967. [Google Scholar] [CrossRef]
  4. Khashan, S.A.; Furlani, E.P. Scalability analysis of magnetic bead separation in a microchannel with an array of soft magnetic elements in a uniform magnetic field. Sep. Purif. Technol. 2014125, 311–318. [Google Scholar] [CrossRef]
  5. Furlani, E.P. Magnetic biotransport: Analysis and applications. Materials 20103, 2412–2446. [Google Scholar] [CrossRef]
  6. Gómez-Pastora, J.; Bringas, E.; Ortiz, I. Design of novel adsorption processes for the removal of arsenic from polluted groundwater employing functionalized magnetic nanoparticles. Chem. Eng. Trans. 201647, 241–246. [Google Scholar]
  7. Gómez-Pastora, J.; Bringas, E.; Lázaro-Díez, M.; Ramos-Vivas, J.; Ortiz, I. The reverse of controlled release: Controlled sequestration of species and biotoxins into nanoparticles (NPs). In Drug Delivery Systems; Stroeve, P., Mahmoudi, M., Eds.; World Scientific: Hackensack, NJ, USA, 2017; pp. 207–244. ISBN 9789813201057. [Google Scholar]
  8. Ruffert, C. Magnetic bead-magic bullet. Micromachines 20167, 21. [Google Scholar] [CrossRef]
  9. Yáñez-Sedeño, P.; Campuzano, S.; Pingarrón, J.M. Magnetic particles coupled to disposable screen printed transducers for electrochemical biosensing. Sensors 201616, 1585. [Google Scholar] [CrossRef]
  10. Schrittwieser, S.; Pelaz, B.; Parak, W.J.; Lentijo-Mozo, S.; Soulantica, K.; Dieckhoff, J.; Ludwig, F.; Guenther, A.; Tschöpe, A.; Schotter, J. Homogeneous biosensing based on magnetic particle labels. Sensors 201616, 828. [Google Scholar] [CrossRef]
  11. He, J.; Huang, M.; Wang, D.; Zhang, Z.; Li, G. Magnetic separation techniques in sample preparation for biological analysis: A review. J. Pharm. Biomed. Anal. 2014101, 84–101. [Google Scholar] [CrossRef]
  12. Ha, Y.; Ko, S.; Kim, I.; Huang, Y.; Mohanty, K.; Huh, C.; Maynard, J.A. Recent advances incorporating superparamagnetic nanoparticles into immunoassays. ACS Appl. Nano Mater. 20181, 512–521. [Google Scholar] [CrossRef]
  13. Gómez-Pastora, J.; González-Fernández, C.; Fallanza, M.; Bringas, E.; Ortiz, I. Flow patterns and mass transfer performance of miscible liquid-liquid flows in various microchannels: Numerical and experimental studies. Chem. Eng. J. 2018344, 487–497. [Google Scholar] [CrossRef]
  14. Gale, B.K.; Jafek, A.R.; Lambert, C.J.; Goenner, B.L.; Moghimifam, H.; Nze, U.C.; Kamarapu, S.K. A review of current methods in microfluidic device fabrication and future commercialization prospects. Inventions 20183, 60. [Google Scholar] [CrossRef]
  15. Nanobiotechnology; Concepts, Applications and Perspectives; Niemeyer, C.M.; Mirkin, C.A. (Eds.) Wiley-VCH: Weinheim, Germany, 2004; ISBN 3527305068. [Google Scholar]
  16. Khashan, S.A.; Dagher, S.; Alazzam, A.; Mathew, B.; Hilal-Alnaqbi, A. Microdevice for continuous flow magnetic separation for bioengineering applications. J. Micromech. Microeng. 201727, 055016. [Google Scholar] [CrossRef]
  17. Basauri, A.; Gomez-Pastora, J.; Fallanza, M.; Bringas, E.; Ortiz, I. Predictive model for the design of reactive micro-separations. Sep. Purif. Technol. 2019209, 900–907. [Google Scholar] [CrossRef]
  18. Abdollahi, P.; Karimi-Sabet, J.; Moosavian, M.A.; Amini, Y. Microfluidic solvent extraction of calcium: Modeling and optimization of the process variables. Sep. Purif. Technol. 2020231, 115875. [Google Scholar] [CrossRef]
  19. Khashan, S.A.; Alazzam, A.; Furlani, E. A novel design for a microfluidic magnetophoresis system: Computational study. In Proceedings of the 12th International Symposium on Fluid Control, Measurement and Visualization (FLUCOME2013), Nara, Japan, 18–23 November 2013. [Google Scholar]
  20. Pamme, N. Magnetism and microfluidics. Lab Chip 20066, 24–38. [Google Scholar] [CrossRef]
  21. Gómez-Pastora, J.; Amiri Roodan, V.; Karampelas, I.H.; Alorabi, A.Q.; Tarn, M.D.; Iles, A.; Bringas, E.; Paunov, V.N.; Pamme, N.; Furlani, E.P.; et al. Two-step numerical approach to predict ferrofluid droplet generation and manipulation inside multilaminar flow chambers. J. Phys. Chem. C 2019123, 10065–10080. [Google Scholar] [CrossRef]
  22. Gómez-Pastora, J.; Karampelas, I.H.; Bringas, E.; Furlani, E.P.; Ortiz, I. Numerical analysis of bead magnetophoresis from flowing blood in a continuous-flow microchannel: Implications to the bead-fluid interactions. Sci. Rep. 20199, 7265. [Google Scholar] [CrossRef]
  23. Tarn, M.D.; Pamme, N. On-Chip Magnetic Particle-Based Immunoassays Using Multilaminar Flow for Clinical Diagnostics. In Microchip Diagnostics Methods and Protocols; Taly, V., Viovy, J.L., Descroix, S., Eds.; Humana Press: New York, NY, USA, 2017; pp. 69–83. [Google Scholar]
  24. Phurimsak, C.; Tarn, M.D.; Peyman, S.A.; Greenman, J.; Pamme, N. On-chip determination of c-reactive protein using magnetic particles in continuous flow. Anal. Chem. 201486, 10552–10559. [Google Scholar] [CrossRef]
  25. Wu, X.; Wu, H.; Hu, Y. Enhancement of separation efficiency on continuous magnetophoresis by utilizing L/T-shaped microchannels. Microfluid. Nanofluid. 201111, 11–24. [Google Scholar] [CrossRef]
  26. Vojtíšek, M.; Tarn, M.D.; Hirota, N.; Pamme, N. Microfluidic devices in superconducting magnets: On-chip free-flow diamagnetophoresis of polymer particles and bubbles. Microfluid. Nanofluid. 201213, 625–635. [Google Scholar] [CrossRef]
  27. Gómez-Pastora, J.; González-Fernández, C.; Real, E.; Iles, A.; Bringas, E.; Furlani, E.P.; Ortiz, I. Computational modeling and fluorescence microscopy characterization of a two-phase magnetophoretic microsystem for continuous-flow blood detoxification. Lab Chip 201818, 1593–1606. [Google Scholar] [CrossRef] [PubMed]
  28. Forbes, T.P.; Forry, S.P. Microfluidic magnetophoretic separations of immunomagnetically labeled rare mammalian cells. Lab Chip 201212, 1471–1479. [Google Scholar] [CrossRef]
  29. Nandy, K.; Chaudhuri, S.; Ganguly, R.; Puri, I.K. Analytical model for the magnetophoretic capture of magnetic microspheres in microfluidic devices. J. Magn. Magn. Mater. 2008320, 1398–1405. [Google Scholar] [CrossRef]
  30. Plouffe, B.D.; Lewis, L.H.; Murthy, S.K. Computational design optimization for microfluidic magnetophoresis. Biomicrofluidics 20115, 013413. [Google Scholar] [CrossRef] [PubMed]
  31. Hale, C.; Darabi, J. Magnetophoretic-based microfluidic device for DNA isolation. Biomicrofluidics 20148, 044118. [Google Scholar] [CrossRef] [PubMed]
  32. Becker, H.; Gärtner, C. Polymer microfabrication methods for microfluidic analytical applications. Electrophoresis 200021, 12–26. [Google Scholar] [CrossRef]
  33. Pekas, N.; Zhang, Q.; Nannini, M.; Juncker, D. Wet-etching of structures with straight facets and adjustable taper into glass substrates. Lab Chip 201010, 494–498. [Google Scholar] [CrossRef]
  34. Wang, T.; Chen, J.; Zhou, T.; Song, L. Fabricating microstructures on glass for microfluidic chips by glass molding process. Micromachines 20189, 269. [Google Scholar] [CrossRef]
  35. Castaño-Álvarez, M.; Pozo Ayuso, D.F.; García Granda, M.; Fernández-Abedul, M.T.; Rodríguez García, J.; Costa-García, A. Critical points in the fabrication of microfluidic devices on glass substrates. Sens. Actuators B Chem. 2008130, 436–448. [Google Scholar] [CrossRef]
  36. Prakash, S.; Kumar, S. Fabrication of microchannels: A review. Proc. Inst. Mech. Eng. Part B J. Eng. Manuf. 2015229, 1273–1288. [Google Scholar] [CrossRef]
  37. Leester-Schädel, M.; Lorenz, T.; Jürgens, F.; Ritcher, C. Fabrication of Microfluidic Devices. In Microsystems for Pharmatechnology: Manipulation of Fluids, Particles, Droplets, and Cells; Dietzel, A., Ed.; Springer: Basel, Switzerland, 2016; pp. 23–57. ISBN 9783319269207. [Google Scholar]
  38. Bartlett, N.W.; Wood, R.J. Comparative analysis of fabrication methods for achieving rounded microchannels in PDMS. J. Micromech. Microeng. 201626, 115013. [Google Scholar] [CrossRef]
  39. Ng, P.F.; Lee, K.I.; Yang, M.; Fei, B. Fabrication of 3D PDMS microchannels of adjustable cross-sections via versatile gel templates. Polymers 201911, 64. [Google Scholar] [CrossRef] [PubMed]
  40. Furlani, E.P.; Sahoo, Y.; Ng, K.C.; Wortman, J.C.; Monk, T.E. A model for predicting magnetic particle capture in a microfluidic bioseparator. Biomed. Microdevices 20079, 451–463. [Google Scholar] [CrossRef]
  41. Tarn, M.D.; Peyman, S.A.; Robert, D.; Iles, A.; Wilhelm, C.; Pamme, N. The importance of particle type selection and temperature control for on-chip free-flow magnetophoresis. J. Magn. Magn. Mater. 2009321, 4115–4122. [Google Scholar] [CrossRef]
  42. Furlani, E.P. Permanent Magnet and Electromechanical Devices; Materials, Analysis and Applications; Academic Press: Waltham, MA, USA, 2001. [Google Scholar]
  43. White, F.M. Viscous Fluid Flow; McGraw-Hill: New York, NY, USA, 1974. [Google Scholar]
  44. Mathew, B.; Alazzam, A.; El-Khasawneh, B.; Maalouf, M.; Destgeer, G.; Sung, H.J. Model for tracing the path of microparticles in continuous flow microfluidic devices for 2D focusing via standing acoustic waves. Sep. Purif. Technol. 2015153, 99–107. [Google Scholar] [CrossRef]
  45. Furlani, E.J.; Furlani, E.P. A model for predicting magnetic targeting of multifunctional particles in the microvasculature. J. Magn. Magn. Mater. 2007312, 187–193. [Google Scholar] [CrossRef]
  46. Furlani, E.P.; Ng, K.C. Analytical model of magnetic nanoparticle transport and capture in the microvasculature. Phys. Rev. E 200673, 061919. [Google Scholar] [CrossRef]
  47. Eibl, R.; Eibl, D.; Pörtner, R.; Catapano, G.; Czermak, P. Cell and Tissue Reaction Engineering; Springer: Berlin/Heidelberg, Germany, 2009. [Google Scholar]
  48. Pamme, N.; Eijkel, J.C.T.; Manz, A. On-chip free-flow magnetophoresis: Separation and detection of mixtures of magnetic particles in continuous flow. J. Magn. Magn. Mater. 2006307, 237–244. [Google Scholar] [CrossRef]
  49. Alorabi, A.Q.; Tarn, M.D.; Gómez-Pastora, J.; Bringas, E.; Ortiz, I.; Paunov, V.N.; Pamme, N. On-chip polyelectrolyte coating onto magnetic droplets-Towards continuous flow assembly of drug delivery capsules. Lab Chip 201717, 3785–3795. [Google Scholar] [CrossRef]
  50. Zhang, H.; Guo, H.; Chen, Z.; Zhang, G.; Li, Z. Application of PECVD SiC in glass micromachining. J. Micromech. Microeng. 200717, 775–780. [Google Scholar] [CrossRef]
  51. Mourzina, Y.; Steffen, A.; Offenhäusser, A. The evaporated metal masks for chemical glass etching for BioMEMS. Microsyst. Technol. 200511, 135–140. [Google Scholar] [CrossRef]
  52. Mata, A.; Fleischman, A.J.; Roy, S. Fabrication of multi-layer SU-8 microstructures. J. Micromech. Microeng. 200616, 276–284. [Google Scholar] [CrossRef]
  53. Su, N. 8 2000 Negative Tone Photoresist Formulations 2002–2025; MicroChem Corporation: Newton, MA, USA, 2002. [Google Scholar]
  54. Su, N. 8 2000 Negative Tone Photoresist Formulations 2035–2100; MicroChem Corporation: Newton, MA, USA, 2002. [Google Scholar]
  55. Fu, C.; Hung, C.; Huang, H. A novel and simple fabrication method of embedded SU-8 micro channels by direct UV lithography. J. Phys. Conf. Ser. 200634, 330–335. [Google Scholar] [CrossRef]
  56. Kazoe, Y.; Yamashiro, I.; Mawatari, K.; Kitamori, T. High-pressure acceleration of nanoliter droplets in the gas phase in a microchannel. Micromachines 20167, 142. [Google Scholar] [CrossRef]
  57. Sharp, K.V.; Adrian, R.J.; Santiago, J.G.; Molho, J.I. Liquid flows in microchannels. In MEMS: Introduction and Fundamentals; Gad-el-Hak, M., Ed.; CRC Press: Boca Raton, FL, USA, 2006; pp. 10-1–10-46. ISBN 9781420036572. [Google Scholar]
  58. Oh, K.W.; Lee, K.; Ahn, B.; Furlani, E.P. Design of pressure-driven microfluidic networks using electric circuit analogy. Lab Chip 201212, 515–545. [Google Scholar] [CrossRef]
  59. Bruus, H. Theoretical Microfluidics; Oxford University Press: New York, NY, USA, 2008; ISBN 9788578110796. [Google Scholar]
  60. Beebe, D.J.; Mensing, G.A.; Walker, G.M. Physics and applications of microfluidics in biology. Annu. Rev. Biomed. Eng. 20024, 261–286. [Google Scholar] [CrossRef] [PubMed]
  61. Yalikun, Y.; Tanaka, Y. Large-scale integration of all-glass valves on a microfluidic device. Micromachines 20167, 83. [Google Scholar] [CrossRef] [PubMed]
  62. Van Heeren, H.; Verhoeven, D.; Atkins, T.; Tzannis, A.; Becker, H.; Beusink, W.; Chen, P. Design Guideline for Microfluidic Device and Component Interfaces (Part 2), Version 3; Available online: http://www.makefluidics.com/en/design-guideline?id=7 (accessed on 9 March 2020).
  63. Scheuble, N.; Iles, A.; Wootton, R.C.R.; Windhab, E.J.; Fischer, P.; Elvira, K.S. Microfluidic technique for the simultaneous quantification of emulsion instabilities and lipid digestion kinetics. Anal. Chem. 201789, 9116–9123. [Google Scholar] [CrossRef] [PubMed]
  64. Lynch, E.C. Red blood cell damage by shear stress. Biophys. J. 197212, 257–273. [Google Scholar]
  65. Paul, R.; Apel, J.; Klaus, S.; Schügner, F.; Schwindke, P.; Reul, H. Shear stress related blood damage in laminar Couette flow. Artif. Organs 200327, 517–529. [Google Scholar] [CrossRef] [PubMed]
  66. Gómez-Pastora, J.; Karampelas, I.H.; Xue, X.; Bringas, E.; Furlani, E.P.; Ortiz, I. Magnetic bead separation from flowing blood in a two-phase continuous-flow magnetophoretic microdevice: Theoretical analysis through computational fluid dynamics simulation. J. Phys. Chem. C 2017121, 7466–7477. [Google Scholar] [CrossRef]
  67. Lim, J.; Yeap, S.P.; Leow, C.H.; Toh, P.Y.; Low, S.C. Magnetophoresis of iron oxide nanoparticles at low field gradient: The role of shape anisotropy. J. Colloid Interface Sci. 2014421, 170–177. [Google Scholar] [CrossRef] [PubMed]
  68. Culbertson, C.T.; Sibbitts, J.; Sellens, K.; Jia, S. Fabrication of Glass Microfluidic Devices. In Microfluidic Electrophoresis: Methods and Protocols; Dutta, D., Ed.; Humana Press: New York, NY, USA, 2019; pp. 1–12. ISBN 978-1-4939-8963-8. [Google Scholar]
Modeling of contactless bubble–bubble interactions in microchannels with integrated inertial pumps

Modeling of contactless bubble–bubble interactions in microchannels with integrated inertial pumps

통합 관성 펌프를 사용하여 마이크로 채널에서 비접촉식 기포-기포 상호 작용 모델링

Physics of Fluids 33, 042002 (2021); https://doi.org/10.1063/5.0041924 B. Hayesa) G. L. Whitingb), and  R. MacCurdyc)

ABSTRACT

In this study, the nonlinear effect of contactless bubble–bubble interactions in inertial micropumps is characterized via reduced parameter one-dimensional and three-dimensional computational fluid dynamics (3D CFD) modeling. A one-dimensional pump model is developed to account for contactless bubble-bubble interactions, and the accuracy of the developed one-dimensional model is assessed via the commercial volume of fluid CFD software, FLOW-3D. The FLOW-3D CFD model is validated against experimental bubble dynamics images as well as experimental pump data. Precollapse and postcollapse bubble and flow dynamics for two resistors in a channel have been successfully explained by the modified one-dimensional model. The net pumping effect design space is characterized as a function of resistor placement and firing time delay. The one-dimensional model accurately predicts cumulative flow for simultaneous resistor firing with inner-channel resistor placements (0.2L < x < 0.8L where L is the channel length) as well as delayed resistor firing with inner-channel resistor placements when the time delay is greater than the time required for the vapor bubble to fill the channel cross section. In general, one-dimensional model accuracy suffers at near-reservoir resistor placements and short time delays which we propose is a result of 3D bubble-reservoir interactions and transverse bubble growth interactions, respectively, that are not captured by the one-dimensional model. We find that the one-dimensional model accuracy improves for smaller channel heights. We envision the developed one-dimensional model as a first-order rapid design tool for inertial pump-based microfluidic systems operating in the contactless bubble–bubble interaction nonlinear regime

이 연구에서 관성 마이크로 펌프에서 비접촉 기포-기포 상호 작용의 비선형 효과는 감소 된 매개 변수 1 차원 및 3 차원 전산 유체 역학 (3D CFD) 모델링을 통해 특성화됩니다. 비접촉식 기포-버블 상호 작용을 설명하기 위해 1 차원 펌프 모델이 개발되었으며, 개발 된 1 차원 모델의 정확도는 유체 CFD 소프트웨어 인 FLOW-3D의 상용 볼륨을 통해 평가됩니다.

FLOW-3D CFD 모델은 실험적인 거품 역학 이미지와 실험적인 펌프 데이터에 대해 검증되었습니다. 채널에 있는 두 저항기의 붕괴 전 및 붕괴 후 기포 및 유동 역학은 수정 된 1 차원 모델에 의해 성공적으로 설명되었습니다. 순 펌핑 효과 설계 공간은 저항 배치 및 발사 시간 지연의 기능으로 특징 지어집니다.

1 차원 모델은 내부 채널 저항 배치 (0.2L <x <0.8L, 여기서 L은 채널 길이)로 동시 저항 발생에 대한 누적 흐름과 시간 지연시 내부 채널 저항 배치로 지연된 저항 발생을 정확하게 예측합니다. 증기 방울이 채널 단면을 채우는 데 필요한 시간보다 큽니다.

일반적으로 1 차원 모델 정확도는 저수지 근처의 저항 배치와 1 차원 모델에 의해 포착되지 않는 3D 기포-저수지 상호 작용 및 가로 기포 성장 상호 작용의 결과 인 짧은 시간 지연에서 어려움을 겪습니다. 채널 높이가 작을수록 1 차원 모델 정확도가 향상됩니다. 우리는 개발 된 1 차원 모델을 비접촉 기포-기포 상호 작용 비선형 영역에서 작동하는 관성 펌프 기반 미세 유체 시스템을 위한 1 차 빠른 설계 도구로 생각합니다.

REFERENCES

1.S. Hassan and X. Zhang, “ Design and fabrication of capillary-driven flow device for point-of-care diagnostics,” Biosensors 10, 39 (2020). https://doi.org/10.3390/bios10040039, Google ScholarCrossref
2.Q. Shizhi and H. Bau, “ Magneto-hydrodynamics based microfluidics,” Mech. Res. Commun. 36, 10 (2009). https://doi.org/10.1016/j.mechrescom.2008.06.013, Google ScholarCrossref
3.N. Mishchuk, T. Heldal, T. Volden, J. Auerswald, and H. Knapp, “ Micropump based on electroosmosis of the second kind,” Electrophoresis 30, 3499 (2009). https://doi.org/10.1002/elps.200900271, Google ScholarCrossref
4.J. Snyder, J. Getpreecharsawas, D. Fang, T. Gaborski, C. Striemer, P. Fauchet, D. Borkholder, and J. McGrath, “ High-performance, low-voltage electroosmotic pumps with molecularly thin silicon nanomembranes,” Proc. Nat. Acad. Sci. U. S. A. 110, 18425–18430 (2013). https://doi.org/10.1073/pnas.1308109110, Google ScholarCrossref
5.K. Vinayakumar, G. Nadiger, V. Shetty, S. Dinesh, M. Nayak, and K. Rajanna, “ Packaged peristaltic micropump for controlled drug delivery application,” Rev. Sci. Instrum. 88, 015102 (2017). https://doi.org/10.1063/1.4973513, Google ScholarScitation, ISI
6.D. Duffy, H. Gillis, J. Lin, N. Sheppard, and G. Kellogg, “ Microfabricated centrifugal microfluidic systems: Characterization and multiple enzymatic assays,” Anal. Chem. 71, 4669 (1999). https://doi.org/10.1021/ac990682c, Google ScholarCrossref
7.V. Gnyawali, M. Saremi, M. Kolios, and S. Tsai, “ Stable microfluidic flow focusing using hydrostatics,” Biomicrofluidics 11, 034104 (2017). https://doi.org/10.1063/1.4983147, Google ScholarScitation, ISI
8.J. Lake, K. Heyde, and W. Ruder, “ Low-cost feedback-controlled syringe pressure pumps for microfluidics applications,” PLoS One 12, e0175089 (2017). https://doi.org/10.1371/journal.pone.0175089, Google ScholarCrossref
9.M. I. Mohammed, S. Haswell, and I. Gibson, “ Lab-on-a-chip or chip-in-a-lab: Challenges of commercialization lost in translation,” Procedia Technology 20, 54–59 (2015), proceedings of The 1st International Design Technology Conference, DESTECH2015, Geelong. Google ScholarCrossref
10.E. Torniainen, A. Govyadinov, D. Markel, and P. Kornilovitch, “ Bubble-driven inertial micropump,” Phys. Fluids 24, 122003 (2012). https://doi.org/10.1063/1.4769755, Google ScholarScitation, ISI
11.H. Hoefemann, S. Wadle, N. Bakhtina, V. Kondrashov, N. Wangler, and R. Zengerle, “ Sorting and lysis of single cells by bubblejet technology,” Sens. Actuators, B 168, 442–445 (2012). https://doi.org/10.1016/j.snb.2012.04.005, Google ScholarCrossref
12.B. Hayes, A. Hayes, M. Rolleston, A. Ferreira, and J. Kirsher, “ Pulsatory mixing of laminar flow using bubble-driven micro-pumps,” in Proceedings of the ASME 2018 International Mechanical Engineering Congress and Exposition (2018), Vol. 7. Google ScholarCrossref
13.E. Ory, H. Yuan, A. Prosperetti, S. Popinet, and S. Zaleski, “ Growth and collapse of a vapor bubble in a narrow tube,” Phys. Fluids 12, 1268 (2000). https://doi.org/10.1063/1.870381, Google ScholarScitation, ISI
14.Z. Yin and A. Prosperetti, “‘ Blinking bubble’ micropump with microfabricated heaters,” J. Micromech. Microeng. 15, 1683 (2005). https://doi.org/10.1088/0960-1317/15/9/010, Google ScholarCrossref
15.M. Einat and M. Grajower, “ Microboiling measurements of thermal-inkjet heaters,” J. Microelectromech. Syst. 19, 391 (2010). https://doi.org/10.1109/JMEMS.2010.2040946, Google ScholarCrossref
16.A. Govyadinov, P. Kornilovitch, D. Markel, and E. Torniainen, “ Single-pulse dynamics and flow rates of inertial micropumps,” Microfluid. Nanofluid. 20, 73 (2016). https://doi.org/10.1007/s10404-016-1738-x, Google ScholarCrossref
17.E. Sourtiji and Y. Peles, “ A micro-synthetic jet in a microchannel using bubble growth and collapse,” Appl. Therm. Eng. 160, 114084 (2019). https://doi.org/10.1016/j.applthermaleng.2019.114084, Google ScholarCrossref
18.B. Hayes, A. Govyadinov, and P. Kornilovitch, “ Microfluidic switchboards with integrated inertial pumps,” Microfluid. Nanofluid. 22, 15 (2018). https://doi.org/10.1007/s10404-017-2032-2, Google ScholarCrossref
19.P. Kornilovitch, A. Govyadinov, D. Markel, and E. Torniainen, “ One-dimensional model of inertial pumping,” Phys. Rev. E 87, 023012 (2013). https://doi.org/10.1103/PhysRevE.87.023012, Google ScholarCrossref
20.H. Yuan and A. Prosperetti, “ The pumping effect of growing and collapsing bubbles in a tube,” J. Micromech. Microeng. 9, 402–413 (1999). https://doi.org/10.1088/0960-1317/9/4/318, Google ScholarCrossref
21.J. Zou, B. Li, and C. Ji, “ Interactions between two oscillating bubbles in a rigid tube,” Exp. Therm. Fluid Sci. 61, 105 (2015). https://doi.org/10.1016/j.expthermflusci.2014.10.021, Google ScholarCrossref
22.C. Hirt and B. Nichols, “ Volume of fluid (vof) method for the dynamics of free boundaries,” J. Comput. Phys. 39, 201–225 (1981). https://doi.org/10.1016/0021-9991(81)90145-5, Google ScholarCrossref
23.C. Borgnakke and R. E. Sonntag, Fundamentals of Thermodynamics, 8th ed. ( Wiley, 1999). Google Scholar
24.O. E. Ruiz, “ CFD model of the thermal inkjet droplet ejection process,” in Proceeding of Heat Transfer Summer Conference (2007), Vol. 3. Google ScholarCrossref
25.T. Theofanous, L. Biasi, H. Isbin, and H. Fauske, “ A theoretical study on bubble growth in constant and time-dependent pressure fields,” Chem. Eng. Sci. 24, 885–897 (1969). https://doi.org/10.1016/0009-2509(69)85008-6, Google ScholarCrossref
26.S. Timoshenko and J. Goodier, Theory of Elasticity, 3rd ed. ( McGaw-Hill, Inc., 1970). Google Scholar

Figure 4. Calculate and simulate the injection of water in a single-channel injection chamber with a nozzle diameter of 60 μm and a thickness of 50 μm, at an operating frequency of 5 KHz, in the X-Y two-dimensional cross-sectional view, at 10, 20, 30, 40 and 200 μs.

DNA Printing Integrated Multiplexer Driver Microelectronic Mechanical System Head (IDMH) and Microfluidic Flow Estimation

DNA 프린팅 통합 멀티플렉서 드라이버 Microelectronic Mechanical System Head (IDMH) 및 Microfluidic Flow Estimation

by Jian-Chiun Liou 1,*,Chih-Wei Peng 1,Philippe Basset 2 andZhen-Xi Chen 11School of Biomedical Engineering, Taipei Medical University, Taipei 11031, Taiwan2ESYCOM, Université Gustave Eiffel, CNRS, CNAM, ESIEE Paris, F-77454 Marne-la-Vallée, France*Author to whom correspondence should be addressed.

Abstract

The system designed in this study involves a three-dimensional (3D) microelectronic mechanical system chip structure using DNA printing technology. We employed diverse diameters and cavity thickness for the heater. DNA beads were placed in this rapid array, and the spray flow rate was assessed. Because DNA cannot be obtained easily, rapidly deploying DNA while estimating the total amount of DNA being sprayed is imperative. DNA printings were collected in a multiplexer driver microelectronic mechanical system head, and microflow estimation was conducted. Flow-3D was used to simulate the internal flow field and flow distribution of the 3D spray room. The simulation was used to calculate the time and pressure required to generate heat bubbles as well as the corresponding mean outlet speed of the fluid. The “outlet speed status” function in Flow-3D was used as a power source for simulating the ejection of fluid by the chip nozzle. The actual chip generation process was measured, and the starting voltage curve was analyzed. Finally, experiments on flow rate were conducted, and the results were discussed. The density of the injection nozzle was 50, the size of the heater was 105 μm × 105 μm, and the size of the injection nozzle hole was 80 μm. The maximum flow rate was limited to approximately 3.5 cc. The maximum flow rate per minute required a power between 3.5 W and 4.5 W. The number of injection nozzles was multiplied by 100. On chips with enlarged injection nozzle density, experiments were conducted under a fixed driving voltage of 25 V. The flow curve obtained from various pulse widths and operating frequencies was observed. The operating frequency was 2 KHz, and the pulse width was 4 μs. At a pulse width of 5 μs and within the power range of 4.3–5.7 W, the monomer was injected at a flow rate of 5.5 cc/min. The results of this study may be applied to estimate the flow rate and the total amount of the ejection liquid of a DNA liquid.

이 연구에서 설계된 시스템은 DNA 프린팅 기술을 사용하는 3 차원 (3D) 마이크로 전자 기계 시스템 칩 구조를 포함합니다. 히터에는 다양한 직경과 캐비티 두께를 사용했습니다. DNA 비드를 빠른 어레이에 배치하고 스프레이 유속을 평가했습니다.

DNA를 쉽게 얻을 수 없기 때문에 DNA를 빠르게 배치하면서 스프레이 되는 총 DNA 양을 추정하는 것이 필수적입니다. DNA 프린팅은 멀티플렉서 드라이버 마이크로 전자 기계 시스템 헤드에 수집되었고 마이크로 플로우 추정이 수행되었습니다.

Flow-3D는 3D 스프레이 룸의 내부 유동장과 유동 분포를 시뮬레이션 하는데 사용되었습니다. 시뮬레이션은 열 거품을 생성하는데 필요한 시간과 압력뿐만 아니라 유체의 해당 평균 출구 속도를 계산하는데 사용되었습니다.

Flow-3D의 “출구 속도 상태”기능은 칩 노즐에 의한 유체 배출 시뮬레이션을 위한 전원으로 사용되었습니다. 실제 칩 생성 프로세스를 측정하고 시작 전압 곡선을 분석했습니다. 마지막으로 유속 실험을 하고 그 결과를 논의했습니다. 분사 노즐의 밀도는 50, 히터의 크기는 105μm × 105μm, 분사 노즐 구멍의 크기는 80μm였다. 최대 유량은 약 3.5cc로 제한되었습니다. 분당 최대 유량은 3.5W에서 4.5W 사이의 전력이 필요했습니다. 분사 노즐의 수에 100을 곱했습니다. 분사 노즐 밀도가 확대 된 칩에 대해 25V의 고정 구동 전압에서 실험을 수행했습니다. 얻은 유동 곡선 다양한 펄스 폭과 작동 주파수에서 관찰되었습니다. 작동 주파수는 2KHz이고 펄스 폭은 4μs입니다. 5μs의 펄스 폭과 4.3–5.7W의 전력 범위 내에서 단량체는 5.5cc / min의 유속으로 주입되었습니다. 이 연구의 결과는 DNA 액체의 토 출액의 유량과 총량을 추정하는 데 적용될 수 있습니다.

Keywords: DNA printingflow estimationMEMS

Introduction

잉크젯 프린트 헤드 기술은 매우 중요하며, 잉크젯 기술의 거대한 발전은 주로 잉크젯 프린트 헤드 기술의 원리 개발에서 시작되었습니다. 잉크젯 인쇄 연구를 위한 대규모 액적 생성기 포함 [ 1 , 2 , 3 , 4 , 5 , 6 , 7 , 8]. 연속 식 잉크젯 시스템은 고주파 응답과 고속 인쇄의 장점이 있습니다. 그러나이 방법의 잉크젯 프린트 헤드의 구조는 더 복잡하고 양산이 어려운 가압 장치, 대전 전극, 편향 전계가 필요하다. 주문형 잉크젯 시스템의 잉크젯 프린트 헤드는 구조가 간단하고 잉크젯 헤드의 다중 노즐을 쉽게 구현할 수 있으며 디지털화 및 색상 지정이 쉽고 이미지 품질은 비교적 좋지만 일반적인 잉크 방울 토출 속도는 낮음 [ 9 , 10 , 11 ].

핫 버블 잉크젯 헤드의 총 노즐 수는 수백 또는 수천에 달할 수 있습니다. 노즐은 매우 미세하여 풍부한 조화 색상과 부드러운 메쉬 톤을 생성할 수 있습니다. 잉크 카트리지와 노즐이 일체형 구조를 이루고 있으며, 잉크 카트리지 교체시 잉크젯 헤드가 동시에 업데이트되므로 노즐 막힘에 대한 걱정은 없지만 소모품 낭비가 발생하고 상대적으로 높음 비용. 주문형 잉크젯 기술은 배출해야 하는 그래픽 및 텍스트 부분에만 잉크 방울을 배출하고 빈 영역에는 잉크 방울이 배출되지 않습니다. 이 분사 방법은 잉크 방울을 충전할 필요가 없으며 전극 및 편향 전기장을 충전할 필요도 없습니다. 노즐 구조가 간단하고 노즐의 멀티 노즐 구현이 용이하며, 출력 품질이 더욱 개선되었습니다. 펄스 제어를 통해 디지털화가 쉽습니다. 그러나 잉크 방울의 토출 속도는 일반적으로 낮습니다. 열 거품 잉크젯, 압전 잉크젯 및 정전기 잉크젯의 세 가지 일반적인 유형이 있습니다. 물론 다른 유형이 있습니다.

압전 잉크젯 기술의 실현 원리는 인쇄 헤드의 노즐 근처에 많은 소형 압전 세라믹을 배치하면 압전 크리스탈이 전기장의 작용으로 변형됩니다. 잉크 캐비티에서 돌출되어 노즐에서 분사되는 패턴 데이터 신호는 압전 크리스탈의 변형을 제어한 다음 잉크 분사량을 제어합니다. 압전 MEMS 프린트 헤드를 사용한 주문형 드롭 하이브리드 인쇄 [ 12]. 열 거품 잉크젯 기술의 실현 원리는 가열 펄스 (기록 신호)의 작용으로 노즐의 발열체 온도가 상승하여 근처의 잉크 용매가 증발하여 많은 수의 핵 형성 작은 거품을 생성하는 것입니다. 내부 거품의 부피는 계속 증가합니다. 일정 수준에 도달하면 생성된 압력으로 인해 잉크가 노즐에서 분사되고 최종적으로 기판 표면에 도달하여 패턴 정보가 재생됩니다 [ 13 , 14 , 15 , 16 , 17 , 18 ].

“3D 제품 프린팅”및 “증분 빠른 제조”의 의미는 진화했으며 모든 증분 제품 제조 기술을 나타냅니다. 이는 이전 제작과는 다른 의미를 가지고 있지만, 자동 제어 하에 소재를 쌓아 올리는 3D 작업 제작 과정의 공통적 인 특징을 여전히 반영하고 있습니다 [ 19 , 20 , 21 , 22 , 23 , 24 ].

이 개발 시스템은 열 거품 분사 기술입니다. 이 빠른 어레이에 DNA 비드를 배치하고 스프레이 유속을 평가하기 위해 다른 히터 직경과 캐비티 두께를 설계하는 것입니다. DNA 제트 칩의 부스트 회로 시스템은 큰 흐름을 구동하기위한 신호 소스입니다. 목적은 분사되는 DNA 용액의 양과 출력을 조정하는 것입니다. 입력 전압을 더 높은 출력 전압으로 변환해야 하는 경우 부스트 컨버터가 유일한 선택입니다. 부스트 컨버터는 내부 금속 산화물 반도체 전계 효과 트랜지스터 (MOSFET)를 통해 전압을 충전하여 부스트 출력의 목적을 달성하고, MOSFET이 꺼지면 인덕터는 부하 정류를 통해 방전됩니다.

인덕터의 충전과 방전 사이의 변환 프로세스는 인덕터를 통한 전압의 방향을 반대로 한 다음 점차적으로 입력 작동 전압보다 높은 전압을 증가시킵니다. MOSFET의 스위칭 듀티 사이클은 확실히 부스트 비율을 결정합니다. MOSFET의 정격 전류와 부스트 컨버터의 부스트 비율은 부스트 ​​컨버터의 부하 전류의 상한을 결정합니다. MOSFET의 정격 전압은 출력 전압의 상한을 결정합니다. 일부 부스트 컨버터는 정류기와 MOSFET을 통합하여 동기식 정류를 제공합니다. 통합 MOSFET은 정확한 제로 전류 턴 오프를 달성하여 부스트 변압기를 보다 효율적으로 만듭니다. 최대 전력 점 추적 장치를 통해 입력 전력을 실시간으로 모니터링합니다. 입력 전압이 최대 입력 전력 지점에 도달하면 부스트 컨버터가 작동하기 시작하여 부스트 컨버터가 최대 전력 출력 지점으로 유리 기판에 DNA 인쇄를 하는 데 적합합니다. 일정한 온 타임 생성 회로를 통해 온 타임이 온도 및 칩의 코너 각도에 영향을 받지 않아 시스템의 안정성이 향상됩니다.

잉크젯 프린트 헤드에 사용되는 기술은 매우 중요합니다. 잉크젯 기술의 엄청난 발전은 주로 잉크젯 프린팅에 사용되는 대형 액적 이젝터 [ 1 , 2 , 3 , 4 , 5 , 6 , 7 , 8 ]를 포함하여 잉크젯 프린트 헤드 기술의 이론 개발에서 시작되었습니다 . 연속 잉크젯 시스템은 고주파 응답과 고속 인쇄의 장점을 가지고 있습니다. 잉크젯 헤드의 총 노즐 수는 수백 또는 수천에 달할 수 있으며 이러한 노즐은 매우 복잡합니다. 노즐은 풍부하고 조화로운 색상과 부드러운 메쉬 톤을 생성할 수 있습니다 [ 9 , 10 ,11 ]. 잉크젯은 열 거품 잉크젯, 압전 잉크젯 및 정전 식 잉크젯의 세 가지 주요 유형으로 분류할 수 있습니다. 다른 유형도 사용 중입니다. 압전 잉크젯의 기능은 다음과 같습니다. 많은 소형 압전 세라믹이 잉크젯 헤드 노즐 근처에 배치됩니다. 압전 결정은 전기장 아래에서 변형됩니다. 그 후, 잉크는 잉크 캐비티에서 압착되어 노즐에서 배출됩니다. 패턴의 데이터 신호는 압전 결정의 변형을 제어한 다음 분사되는 잉크의 양을 제어합니다. 압전 마이크로 전자 기계 시스템 (MEMS) 잉크젯 헤드는 하이브리드 인쇄에 사용됩니다. [ 12]. 열 버블 잉크젯 기술은 다음과 같이 작동합니다. 가열 펄스 (즉, 기록 신호) 하에서 노즐의 가열 구성 요소의 온도가 상승하여 근처의 잉크 용매를 증발시켜 많은 양의 작은 핵 기포를 생성합니다. 내부 기포의 부피가 지속적으로 증가합니다. 압력이 일정 수준에 도달하면 노즐에서 잉크가 분출되고 잉크가 기판 표면에 도달하여 패턴과 메시지가 표시됩니다 [ 13 , 14 , 15 , 16 , 17 , 18 ].

3 차원 (3D) 제품 프린팅 및 빠른 프로토 타입 기술의 발전에는 모든 빠른 프로토 타입의 생산 기술이 포함됩니다. 래피드 프로토 타입 기술은 기존 생산 방식과는 다르지만 3D 제품 프린팅 생산 과정의 일부 특성을 공유합니다. 구체적으로 자동 제어 [ 19 , 20 , 21 , 22 , 23 , 24 ] 하에서 자재를 쌓아 올립니다 .

이 연구에서 개발된 시스템은 열 기포 방출 기술을 사용했습니다. 이 빠른 어레이에 DNA 비드를 배치하기 위해 히터에 대해 다른 직경과 다른 공동 두께가 사용되었습니다. 그 후, 스프레이 유속을 평가했다. DNA 제트 칩의 부스트 회로 시스템은 큰 흐름을 구동하기위한 신호 소스입니다. 목표는 분사되는 DNA 액체의 양과 출력을 조정하는 것입니다. 입력 전압을 더 높은 출력 전압으로 수정해야하는 경우 승압 컨버터가 유일한 옵션입니다. 승압 컨버터는 내부 금속 산화물 반도체 전계 효과 트랜지스터 (MOSFET)를 충전하여 출력 전압을 증가시킵니다. MOSFET이 꺼지면 부하 정류를 통해 인덕턴스가 방전됩니다. 충전과 방전 사이에서 인덕터를 변경하는 과정은 인덕터를 통과하는 전압의 방향을 변경합니다. 전압은 입력 작동 전압을 초과하는 지점까지 점차적으로 증가합니다. MOSFET 스위치의 듀티 사이클은 부스트 ​​비율을 결정합니다. MOSFET의 승압 컨버터의 정격 전류와 부스트 비율은 승압 컨버터의 부하 전류의 상한을 결정합니다. MOSFET의 정격 전류는 출력 전압의 상한을 결정합니다. 일부 승압 컨버터는 정류기와 MOSFET을 통합하여 동기식 정류를 제공합니다. 통합 MOSFET은 정밀한 제로 전류 셧다운을 실현할 수 있으므로 셋업 컨버터의 효율성을 높일 수 있습니다. 최대 전력 점 추적 장치는 입력 전력을 실시간으로 모니터링하는 데 사용되었습니다. 입력 전압이 최대 입력 전력 지점에 도달하면 승압 컨버터가 작동을 시작합니다. 스텝 업 컨버터는 DNA 프린팅을 위한 최대 전력 출력 포인트가 있는 유리 기판에 사용됩니다.

MEMS Chip Design for Bubble Jet

이 연구는 히터 크기, 히터 번호 및 루프 저항과 같은 특정 매개 변수를 조작하여 5 가지 유형의 액체 배출 챔버 구조를 설계했습니다. 표 1 은 측정 결과를 나열합니다. 이 시스템은 다양한 히터의 루프 저항을 분석했습니다. 100 개 히터 설계를 완료하기 위해 2 세트의 히터를 사용하여 각 단일 회로 시리즈를 통과하기 때문에 100 개의 히터를 설계할 때 총 루프 저항은 히터 50 개의 총 루프 저항보다 하나 더 커야 합니다. 이 연구에서 MEMS 칩에서 기포를 배출하는 과정에서 저항 층의 면저항은 29 Ω / m 2입니다. 따라서 모델 A의 총 루프 저항이 가장 컸습니다. 일반 사이즈 모델 (모델 B1, C, D, E)의 두 배였습니다. 모델 B1, C, D 및 E의 총 루프 저항은 약 29 Ω / m 2 입니다. 표 1 에 따르면 오류 범위는 허용된 설계 값 이내였습니다. 따라서야 연구에서 설계된 각 유형의 단일 칩은 동일한 생산 절차 결과를 가지며 후속 유량 측정에 사용되었습니다.

Table 1. List of resistance measurement of single circuit resistance.
Table 1. List of resistance measurement of single circuit resistance.

DNA를 뿌린 칩의 파워가 정상으로 확인되면 히터 버블의 성장 특성을 테스트하고 검증했습니다. DNA 스프레이 칩의 필름 두께와 필름 품질은 히터의 작동 조건과 스프레이 품질에 영향을 줍니다. 따라서 기포 성장 현상과 그 성장 특성을 이해하면 본 연구에서 DNA 스프레이 칩의 특성과 작동 조건을 명확히 하는 데 도움이 됩니다.

설계된 시스템은 기포 성장 조건을 관찰하기 위해 개방형 액체 공급 방법을 채택했습니다. 이미지 관찰을 위해 발광 다이오드 (LED, Nichia NSPW500GS-K1, 3.1V 백색 LED 5mm)를 사용하는 동기식 플래시 방식을 사용하여 동기식 지연 광원을 생성했습니다. 이 시스템은 또한 전하 결합 장치 (CCD, Flir Grasshopper3 GigE GS3-PGE-50S5C-C)를 사용하여 이미지를 캡처했습니다. 그림 1핵 형성, 성장, 거품 생성에서 소산에 이르는 거품의 과정을 보여줍니다. 이 시스템은 기포의 성장 및 소산 과정을 확인하여 시작 전압을 관찰하는 데 사용할 수 있습니다. 마이크로 채널의 액체 공급 방법은 LED가 깜빡이는 시간을 가장 큰 기포 발생에 필요한 시간 (15μs)으로 설정했습니다. 이 디자인은 부적합한 깜박임 시간으로 인한 잘못된 판단과 거품 이미지 캡처 불가능을 방지합니다.

Figure 1. The system uses CCD to capture images.
Figure 1. The system uses CCD to capture images.

<내용 중략>…….

Table 2. Open pool test starting voltage results.
Table 2. Open pool test starting voltage results.
Figure 2. Serial input parallel output shift registers forms of connection.
Figure 2. Serial input parallel output shift registers forms of connection.
Figure 3. The geometry of the jet cavity. (a) The actual DNA liquid chamber, (b) the three-dimensional view of the microfluidic single channel. A single-channel jet cavity with 60 μm diameter and 50 μm thickness, with an operating frequency of 5 KHz, in (a) three-dimensional side view (b) X-Z two-dimensional cross-sectional view, at 10, 20, 30, 40 and 200 μs injection conditions.
Figure 3. The geometry of the jet cavity. (a) The actual DNA liquid chamber, (b) the three-dimensional view of the microfluidic single channel. A single-channel jet cavity with 60 μm diameter and 50 μm thickness, with an operating frequency of 5 KHz, in (a) three-dimensional side view (b) X-Z two-dimensional cross-sectional view, at 10, 20, 30, 40 and 200 μs injection conditions.
Figure 4. Calculate and simulate the injection of water in a single-channel injection chamber with a nozzle diameter of 60 μm and a thickness of 50 μm, at an operating frequency of 5 KHz, in the X-Y two-dimensional cross-sectional view, at 10, 20, 30, 40 and 200 μs.
Figure 4. Calculate and simulate the injection of water in a single-channel injection chamber with a nozzle diameter of 60 μm and a thickness of 50 μm, at an operating frequency of 5 KHz, in the X-Y two-dimensional cross-sectional view, at 10, 20, 30, 40 and 200 μs.
Figure 5 depicts the calculation results of the 2D X-Z cross section. At 100 μs and 200 μs, the fluid injection orifice did not completely fill the chamber. This may be because the size of the single-channel injection cavity was unsuitable for the highest operating frequency of 10 KHz. Thus, subsequent calculation simulations employed 5 KHz as the reference operating frequency. The calculation simulation results were calculated according to the operating frequency of the impact. Figure 6 illustrates the injection cavity height as 60 μm and 30 μm and reveals the 2D X-Y cross section. At 100 μs and 200 μs, the fluid injection orifice did not completely fill the chamber. In those stages, the fluid was still filling the chamber, and the flow field was not yet stable.
Figure 5 depicts the calculation results of the 2D X-Z cross section. At 100 μs and 200 μs, the fluid injection orifice did not completely fill the chamber. This may be because the size of the single-channel injection cavity was unsuitable for the highest operating frequency of 10 KHz. Thus, subsequent calculation simulations employed 5 KHz as the reference operating frequency. The calculation simulation results were calculated according to the operating frequency of the impact. Figure 6 illustrates the injection cavity height as 60 μm and 30 μm and reveals the 2D X-Y cross section. At 100 μs and 200 μs, the fluid injection orifice did not completely fill the chamber. In those stages, the fluid was still filling the chamber, and the flow field was not yet stable.
Figure 6. Calculate and simulate water in a single-channel spray chamber with a spray hole diameter of 60 μm and a thickness of 50 μm, with an operating frequency of 10 KHz, in an XY cross-sectional view, at 10, 20, 30, 40, 100, 110, 120, 130, 140 and 200 μs injection situation.
Figure 6. Calculate and simulate water in a single-channel spray chamber with a spray hole diameter of 60 μm and a thickness of 50 μm, with an operating frequency of 10 KHz, in an XY cross-sectional view, at 10, 20, 30, 40, 100, 110, 120, 130, 140 and 200 μs injection situation.
Figure 7. The DNA printing integrated multiplexer driver MEMS head (IDMH).
Figure 7. The DNA printing integrated multiplexer driver MEMS head (IDMH).
Figure 8. The initial voltage diagrams of chip number A,B,C,D,E type.
Figure 8. The initial voltage diagrams of chip number A,B,C,D,E type.
Figure 9. The initial energy diagrams of chip number A,B,C,D,E type.
Figure 9. The initial energy diagrams of chip number A,B,C,D,E type.
Figure 10. A Type-Sample01 flow test.
Figure 10. A Type-Sample01 flow test.
Figure 11. A Type-Sample01 drop volume.
Figure 11. A Type-Sample01 drop volume.
Figure 12. A Type-Sample01 flow rate.
Figure 12. A Type-Sample01 flow rate.
Figure 13. B1-00 flow test.
Figure 13. B1-00 flow test.
Figure 14. C Type-01 flow test.
Figure 14. C Type-01 flow test.
Figure 15. D Type-02 flow test.
Figure 15. D Type-02 flow test.
Figure 16. E1 type flow test.
Figure 16. E1 type flow test.
Figure 17. E1 type ejection rate relationship.
Figure 17. E1 type ejection rate relationship.

Conclusions

이 연구는 DNA 프린팅 IDMH를 제공하고 미세 유체 흐름 추정을 수행했습니다. 설계된 DNA 스프레이 캐비티와 20V의 구동 전압에서 다양한 펄스 폭의 유동 성능이 펄스 폭에 따라 증가하는 것으로 밝혀졌습니다.

E1 유형 유량 테스트는 해당 유량이 3.1cc / min으로 증가함에 따라 유량이 전력 변화에 영향을 받는 것으로 나타났습니다. 동력이 증가함에 따라 유량은 0.75cc / min에서 3.5cc / min으로 최대 6.5W까지 증가했습니다. 동력이 더 증가하면 유량은 에너지와 함께 증가하지 않습니다. 이것은 이 테이블 디자인이 가장 크다는 것을 보여줍니다. 유속은 3.5cc / 분이었다.
작동 주파수가 2KHz이고 펄스 폭이 4μs 및 5μs 인 특수 설계된 DNA 스프레이 룸 구조에서 다양한 전력 조건 하에서 유량 변화를 관찰했습니다. 4.3–5.87 W의 출력 범위 내에서 주입 된 모노머의 유속은 5.5cc / 분이었습니다. 이것은 힘이 증가해도 변하지 않았습니다. DNA는 귀중하고 쉽게 얻을 수 없습니다. 이 실험을 통해 우리는 DNA가 뿌려진 마이크로 어레이 바이오칩의 수천 개의 지점에 필요한 총 DNA 양을 정확하게 추정 할 수 있습니다.

<내용 중략>…….

References

  1. Pydar, O.; Paredes, C.; Hwang, Y.; Paz, J.; Shah, N.; Candler, R. Characterization of 3D-printed microfluidic chip interconnects with integrated O-rings. Sens. Actuators Phys. 2014205, 199–203. [Google Scholar] [CrossRef]
  2. Ohtani, K.; Tsuchiya, M.; Sugiyama, H.; Katakura, T.; Hayakawa, M.; Kanai, T. Surface treatment of flow channels in microfluidic devices fabricated by stereolitography. J. Oleo Sci. 201463, 93–96. [Google Scholar] [CrossRef]
  3. Castrejn-Pita, J.R.; Martin, G.D.; Hoath, S.D.; Hutchings, I.M. A simple large-scale droplet generator for studies of inkjet printing. Rev. Sci. Instrum. 200879, 075108. [Google Scholar] [CrossRef] [PubMed]
  4. Asai, A. Application of the nucleation theory to the design of bubble jet printers. Jpn. J. Appl. Phys. Regul. Rap. Short Notes 198928, 909–915. [Google Scholar] [CrossRef]
  5. Aoyama, R.; Seki, M.; Hong, J.W.; Fujii, T.; Endo, I. Novel Liquid Injection Method with Wedge-shaped Microchannel on a PDMS Microchip System for Diagnostic Analyses. In Transducers’ 01 Eurosensors XV; Springer: Berlin, Germany, 2001; pp. 1204–1207. [Google Scholar]
  6. Xu, B.; Zhang, Y.; Xia, H.; Dong, W.; Ding, H.; Sun, H. Fabrication and multifunction integration of microfluidic chips by femtosecond laser direct writing. Lab Chip 201313, 1677–1690. [Google Scholar] [CrossRef] [PubMed]
  7. Nayve, R.; Fujii, M.; Fukugawa, A.; Takeuchi, T.; Murata, M.; Yamada, Y. High-Resolution long-array thermal ink jet printhead fabricated by anisotropic wet etching and deep Si RIE. J. Microelectromech. Syst. 200413, 814–821. [Google Scholar] [CrossRef]
  8. O’Connor, J.; Punch, J.; Jeffers, N.; Stafford, J. A dimensional comparison between embedded 3D: Printed and silicon microchannesl. J. Phys. Conf. Ser. 2014525, 012009. [Google Scholar] [CrossRef]
  9. Fang, Y.J.; Lee, J.I.; Wang, C.H.; Chung, C.K.; Ting, J. Modification of heater and bubble clamping behavior in off-shooting inkjet ejector. In Proceedings of the IEEE Sensors, Irvine, CA, USA, 30 October–3 November 2005; pp. 97–100. [Google Scholar]
  10. Lee, W.; Kwon, D.; Choi, W.; Jung, G.; Jeon, S. 3D-Printed microfluidic device for the detection of pathogenic bacteria using size-based separation in helical channel with trapezoid cross-section. Sci. Rep. 20155, 7717. [Google Scholar] [CrossRef] [PubMed]
  11. Shin, D.Y.; Smith, P.J. Theoretical investigation of the influence of nozzle diameter variation on the fabrication of thin film transistor liquid crystal display color filters. J. Appl. Phys. 2008103, 114905-1–114905-11. [Google Scholar] [CrossRef]
  12. Kim, Y.; Kim, S.; Hwang, J.; Kim, Y. Drop-on-Demand hybrid printing using piezoelectric MEMS printhead at various waveforms, high voltages and jetting frequencies. J. Micromech. Microeng. 201323, 8. [Google Scholar] [CrossRef]
  13. Shin, S.J.; Kuka, K.; Shin, J.W.; Lee, C.S.; Oha, Y.S.; Park, S.O. Thermal design modifications to improve firing frequency of back shooting inkjet printhead. Sens. Actuators Phys. 2004114, 387–391. [Google Scholar] [CrossRef]
  14. Rose, D. Microfluidic Technologies and Instrumentation for Printing DNA Microarrays. In Microarray Biochip Technology; Eaton Publishing: Norwalk, CT, USA, 2000; p. 35. [Google Scholar]
  15. Wu, D.; Wu, S.; Xu, J.; Niu, L.; Midorikawa, K.; Sugioka, K. Hybrid femtosecond laser microfabrication to achieve true 3D glass/polymer composite biochips with multiscale features and high performance: The concept of ship-in-abottle biochip. Laser Photon. Rev. 20148, 458–467. [Google Scholar] [CrossRef]
  16. McIlroy, C.; Harlen, O.; Morrison, N. Modelling the jetting of dilute polymer solutions in drop-on-demand inkjet printing. J. Non Newton. Fluid Mech. 2013201, 17–28. [Google Scholar] [CrossRef]
  17. Anderson, K.; Lockwood, S.; Martin, R.; Spence, D. A 3D printed fluidic device that enables integrated features. Anal. Chem. 201385, 5622–5626. [Google Scholar] [CrossRef] [PubMed]
  18. Avedisian, C.T.; Osborne, W.S.; McLeod, F.D.; Curley, C.M. Measuring bubble nucleation temperature on the surface of a rapidly heated thermal ink-jet heater immersed in a pool of water. Proc. R. Soc. A Lond. Math. Phys. Sci. 1999455, 3875–3899. [Google Scholar] [CrossRef]
  19. Lim, J.H.; Kuk, K.; Shin, S.J.; Baek, S.S.; Kim, Y.J.; Shin, J.W.; Oh, Y.S. Failure mechanisms in thermal inkjet printhead analyzed by experiments and numerical simulation. Microelectron. Reliab. 200545, 473–478. [Google Scholar] [CrossRef]
  20. Shallan, A.; Semjkal, P.; Corban, M.; Gujit, R.; Breadmore, M. Cost-Effective 3D printing of visibly transparent microchips within minutes. Anal. Chem. 201486, 3124–3130. [Google Scholar] [CrossRef] [PubMed]
  21. Cavicchi, R.E.; Avedisian, C.T. Bubble nucleation and growth anomaly for a hydrophilic microheater attributed to metastable nanobubbles. Phys. Rev. Lett. 200798, 124501. [Google Scholar] [CrossRef] [PubMed]
  22. Kamei, K.; Mashimo, Y.; Koyama, Y.; Fockenberg, C.; Nakashima, M.; Nakajima, M.; Li, J.; Chen, Y. 3D printing of soft lithography mold for rapid production of polydimethylsiloxane-based microfluidic devices for cell stimulation with concentration gradients. Biomed. Microdevices 201517, 36. [Google Scholar] [CrossRef] [PubMed]
  23. Shin, S.J.; Kuka, K.; Shin, J.W.; Lee, C.S.; Oha, Y.S.; Park, S.O. Firing frequency improvement of back shooting inkjet printhead by thermal management. In Proceedings of the TRANSDUCERS’03. 12th International Conference on Solid-State Sensors.Actuators and Microsystems. Digest of Technical Papers (Cat. No.03TH8664), Boston, MA, USA, 8–12 June 2003; Volume 1, pp. 380–383. [Google Scholar]
  24. Laio, X.; Song, J.; Li, E.; Luo, Y.; Shen, Y.; Chen, D.; Chen, Y.; Xu, Z.; Sugoioka, K.; Midorikawa, K. Rapid prototyping of 3D microfluidic mixers in glass by femtosecond laser direct writing. Lab Chip 201212, 746–749. [Google Scholar] [CrossRef] [PubMed]
Figure 3. (a) Velocity distribution in a section perpendicular to the flow for rectangular (left) and Ushaped (right) cross section channels, and (b) particle location in these cross sections.

Continuous-Flow Separation of Magnetic Particles from Biofluids: How Does the Microdevice Geometry Determine the Separation Performance?

Cristina González Fernández,1 Jenifer Gómez Pastora,2 Arantza Basauri,1 Marcos Fallanza,1 Eugenio Bringas,1 Jeffrey J. Chalmers,2 and Inmaculada Ortiz1,*
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생체 유체에서 자성 입자의 연속 흐름 분리 : 마이크로 장치 형상이 분리 성능을 어떻게 결정합니까?

Abstract

The use of functionalized magnetic particles for the detection or separation of multiple chemicals and biomolecules from biofluids continues to attract significant attention. After their incubation with the targeted substances, the beads can be magnetically recovered to perform analysis or diagnostic tests. Particle recovery with permanent magnets in continuous-flow microdevices has gathered great attention in the last decade due to the multiple advantages of microfluidics. As such, great efforts have been made to determine the magnetic and fluidic conditions for achieving complete particle capture; however, less attention has been paid to the effect of the channel geometry on the system performance, although it is key for designing systems that simultaneously provide high particle recovery and flow rates. Herein, we address the optimization of Y-Y-shaped microchannels, where magnetic beads are separated from blood and collected into a buffer stream by applying an external magnetic field. The influence of several geometrical features (namely cross section shape, thickness, length, and volume) on both bead recovery and system throughput is studied. For that purpose, we employ an experimentally validated Computational Fluid Dynamics (CFD) numerical model that considers the dominant forces acting on the beads during separation. Our results indicate that rectangular, long devices display the best performance as they deliver high particle recovery and high throughput. Thus, this methodology could be applied to the rational design of lab-on-a-chip devices for any magnetically driven purification, enrichment or isolation.

생체 유체에서 여러 화학 물질과 생체 분자의 검출 또는 분리를 위한 기능화된 자성 입자의 사용은 계속해서 상당한 관심을 받고 있습니다. 표적 물질과 함께 배양 한 후 비드는 자기적으로 회수되어 분석 또는 진단 테스트를 수행 할 수 있습니다.

연속 흐름 마이크로 장치에서 영구 자석을 사용한 입자 회수는 마이크로 유체의 여러 장점으로 인해 지난 10 년 동안 큰 관심을 모았습니다. 따라서 완전한 입자 포획을 달성하기 위한 자기 및 유체 조건을 결정하기 위해 많은 노력을 기울였습니다.

그러나 높은 입자 회수율과 유속을 동시에 제공하는 시스템을 설계하는데 있어 핵심이기는 하지만 시스템 성능에 대한 채널 형상의 영향에 대해서는 덜 주의를 기울였습니다.

여기에서 우리는 자기 비드가 혈액에서 분리되어 외부 자기장을 적용하여 버퍼 스트림으로 수집되는 Y-Y 모양의 마이크로 채널의 최적화를 다룹니다. 비드 회수 및 시스템 처리량에 대한 여러 기하학적 특징 (즉, 단면 형상, 두께, 길이 및 부피)의 영향을 연구합니다.

이를 위해 분리 중에 비드에 작용하는 지배적인 힘을 고려하는 실험적으로 검증된 CFD (Computational Fluid Dynamics) 수치 모델을 사용합니다.

우리의 결과는 직사각형의 긴 장치가 높은 입자 회수율과 높은 처리량을 제공하기 때문에 최고의 성능을 보여줍니다. 따라서 이 방법론은 자기 구동 정제, 농축 또는 분리를 위한 랩 온어 칩 장치의 합리적인 설계에 적용될 수 있습니다.

Keywords: particle magnetophoresis, CFD, cross section, chip fabrication

Figure 1 (a) Top view of the microfluidic-magnetophoretic device, (b) Schematic representation of the channel cross-sections studied in this work, and (c) the magnet position relative to the channel location (Sepy and Sepz are the magnet separation distances in y and z, respectively).
Figure 1 (a) Top view of the microfluidic-magnetophoretic device, (b) Schematic representation of the channel cross-sections studied in this work, and (c) the magnet position relative to the channel location (Sepy and Sepz are the magnet separation distances in y and z, respectively).
Figure 2. (a) Channel-magnet configuration and (b–d) magnetic force distribution in the channel midplane for 2 mm, 5 mm and 10 mm long rectangular (left) and U-shaped (right) devices.
Figure 2. (a) Channel-magnet configuration and (b–d) magnetic force distribution in the channel midplane for 2 mm, 5 mm and 10 mm long rectangular (left) and U-shaped (right) devices.
Figure 3. (a) Velocity distribution in a section perpendicular to the flow for rectangular (left) and Ushaped (right) cross section channels, and (b) particle location in these cross sections.
Figure 3. (a) Velocity distribution in a section perpendicular to the flow for rectangular (left) and Ushaped (right) cross section channels, and (b) particle location in these cross sections.
Figure 4. Influence of fluid flow rate on particle recovery when the applied magnetic force is (a) different and (b) equal in U-shaped and rectangular cross section microdevices.
Figure 4. Influence of fluid flow rate on particle recovery when the applied magnetic force is (a) different and (b) equal in U-shaped and rectangular cross section microdevices.
Figure 5. Magnetic bead capture as a function of fluid flow rate for all of the studied geometries.
Figure 5. Magnetic bead capture as a function of fluid flow rate for all of the studied geometries.
Figure 6. Influence of (a) magnetic and fluidic forces (J parameter) and (b) channel geometry (θ parameter) on particle recovery. Note that U-2mm does not accurately fit a line.
Figure 6. Influence of (a) magnetic and fluidic forces (J parameter) and (b) channel geometry (θ parameter) on particle recovery. Note that U-2mm does not accurately fit a line.
Figure 7. Dependence of bead capture on the (a) functional channel volume, and (b) particle residence time (tres). Note that in the curve fitting expressions V represents the functional channel volume and that U-2mm does not accurately fit a line.
Figure 7. Dependence of bead capture on the (a) functional channel volume, and (b) particle residence time (tres). Note that in the curve fitting expressions V represents the functional channel volume and that U-2mm does not accurately fit a line.

References

  1. Gómez-Pastora J., Xue X., Karampelas I.H., Bringas E., Furlani E.P., Ortiz I. Analysis of separators for magnetic beads recovery: From large systems to multifunctional microdevices. Sep. Purif. Technol. 2017;172:16–31. doi: 10.1016/j.seppur.2016.07.050. [CrossRef] [Google Scholar]
  2. Wise N., Grob T., Morten K., Thompson I., Sheard S. Magnetophoretic velocities of superparamagnetic particles, agglomerates and complexes. J. Magn. Magn. Mater. 2015;384:328–334. doi: 10.1016/j.jmmm.2015.02.031. [CrossRef] [Google Scholar]
  3. Khashan S.A., Elnajjar E., Haik Y. CFD simulation of the magnetophoretic separation in a microchannel. J. Magn. Magn. Mater. 2011;323:2960–2967. doi: 10.1016/j.jmmm.2011.06.001. [CrossRef] [Google Scholar]
  4. Khashan S.A., Furlani E.P. Scalability analysis of magnetic bead separation in a microchannel with an array of soft magnetic elements in a uniform magnetic field. Sep. Purif. Technol. 2014;125:311–318. doi: 10.1016/j.seppur.2014.02.007. [CrossRef] [Google Scholar]
  5. Furlani E.P. Magnetic biotransport: Analysis and applications. Materials. 2010;3:2412–2446. doi: 10.3390/ma3042412. [CrossRef] [Google Scholar]
  6. Gómez-Pastora J., Bringas E., Ortiz I. Design of novel adsorption processes for the removal of arsenic from polluted groundwater employing functionalized magnetic nanoparticles. Chem. Eng. Trans. 2016;47:241–246. [Google Scholar]
  7. Gómez-Pastora J., Bringas E., Lázaro-Díez M., Ramos-Vivas J., Ortiz I. The reverse of controlled release: Controlled sequestration of species and biotoxins into nanoparticles (NPs) In: Stroeve P., Mahmoudi M., editors. Drug Delivery Systems. World Scientific; Hackensack, NJ, USA: 2017. pp. 207–244. [Google Scholar]
  8. Ruffert C. Magnetic bead-magic bullet. Micromachines. 2016;7:21. doi: 10.3390/mi7020021. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  9. Yáñez-Sedeño P., Campuzano S., Pingarrón J.M. Magnetic particles coupled to disposable screen printed transducers for electrochemical biosensing. Sensors. 2016;16:1585. doi: 10.3390/s16101585. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  10. Schrittwieser S., Pelaz B., Parak W.J., Lentijo-Mozo S., Soulantica K., Dieckhoff J., Ludwig F., Guenther A., Tschöpe A., Schotter J. Homogeneous biosensing based on magnetic particle labels. Sensors. 2016;16:828. doi: 10.3390/s16060828. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  11. He J., Huang M., Wang D., Zhang Z., Li G. Magnetic separation techniques in sample preparation for biological analysis: A review. J. Pharm. Biomed. Anal. 2014;101:84–101. doi: 10.1016/j.jpba.2014.04.017. [PubMed] [CrossRef] [Google Scholar]
  12. Ha Y., Ko S., Kim I., Huang Y., Mohanty K., Huh C., Maynard J.A. Recent advances incorporating superparamagnetic nanoparticles into immunoassays. ACS Appl. Nano Mater. 2018;1:512–521. doi: 10.1021/acsanm.7b00025. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  13. Gómez-Pastora J., González-Fernández C., Fallanza M., Bringas E., Ortiz I. Flow patterns and mass transfer performance of miscible liquid-liquid flows in various microchannels: Numerical and experimental studies. Chem. Eng. J. 2018;344:487–497. doi: 10.1016/j.cej.2018.03.110. [CrossRef] [Google Scholar]
  14. Gale B.K., Jafek A.R., Lambert C.J., Goenner B.L., Moghimifam H., Nze U.C., Kamarapu S.K. A review of current methods in microfluidic device fabrication and future commercialization prospects. Inventions. 2018;3:60. doi: 10.3390/inventions3030060. [CrossRef] [Google Scholar]
  15. Niemeyer C.M., Mirkin C.A., editors. Nanobiotechnology; Concepts, Applications and Perspectives. Wiley-VCH; Weinheim, Germany: 2004. [Google Scholar]
  16. Khashan S.A., Dagher S., Alazzam A., Mathew B., Hilal-Alnaqbi A. Microdevice for continuous flow magnetic separation for bioengineering applications. J. Micromech. Microeng. 2017;27:055016. doi: 10.1088/1361-6439/aa666d. [CrossRef] [Google Scholar]
  17. Basauri A., Gomez-Pastora J., Fallanza M., Bringas E., Ortiz I. Predictive model for the design of reactive micro-separations. Sep. Purif. Technol. 2019;209:900–907. doi: 10.1016/j.seppur.2018.09.028. [CrossRef] [Google Scholar]
  18. Abdollahi P., Karimi-Sabet J., Moosavian M.A., Amini Y. Microfluidic solvent extraction of calcium: Modeling and optimization of the process variables. Sep. Purif. Technol. 2020;231:115875. doi: 10.1016/j.seppur.2019.115875. [CrossRef] [Google Scholar]
  19. Khashan S.A., Alazzam A., Furlani E. A novel design for a microfluidic magnetophoresis system: Computational study; Proceedings of the 12th International Symposium on Fluid Control, Measurement and Visualization (FLUCOME2013); Nara, Japan. 18–23 November 2013. [Google Scholar]
  20. Pamme N. Magnetism and microfluidics. Lab Chip. 2006;6:24–38. doi: 10.1039/B513005K. [PubMed] [CrossRef] [Google Scholar]
  21. Gómez-Pastora J., Amiri Roodan V., Karampelas I.H., Alorabi A.Q., Tarn M.D., Iles A., Bringas E., Paunov V.N., Pamme N., Furlani E.P., et al. Two-step numerical approach to predict ferrofluid droplet generation and manipulation inside multilaminar flow chambers. J. Phys. Chem. C. 2019;123:10065–10080. doi: 10.1021/acs.jpcc.9b01393. [CrossRef] [Google Scholar]
  22. Gómez-Pastora J., Karampelas I.H., Bringas E., Furlani E.P., Ortiz I. Numerical analysis of bead magnetophoresis from flowing blood in a continuous-flow microchannel: Implications to the bead-fluid interactions. Sci. Rep. 2019;9:7265. doi: 10.1038/s41598-019-43827-x. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  23. Tarn M.D., Pamme N. On-Chip Magnetic Particle-Based Immunoassays Using Multilaminar Flow for Clinical Diagnostics. In: Taly V., Viovy J.L., Descroix S., editors. Microchip Diagnostics Methods and Protocols. Humana Press; New York, NY, USA: 2017. pp. 69–83. [Google Scholar]
  24. Phurimsak C., Tarn M.D., Peyman S.A., Greenman J., Pamme N. On-chip determination of c-reactive protein using magnetic particles in continuous flow. Anal. Chem. 2014;86:10552–10559. doi: 10.1021/ac5023265. [PubMed] [CrossRef] [Google Scholar]
  25. Wu X., Wu H., Hu Y. Enhancement of separation efficiency on continuous magnetophoresis by utilizing L/T-shaped microchannels. Microfluid. Nanofluid. 2011;11:11–24. doi: 10.1007/s10404-011-0768-7. [CrossRef] [Google Scholar]
  26. Vojtíšek M., Tarn M.D., Hirota N., Pamme N. Microfluidic devices in superconducting magnets: On-chip free-flow diamagnetophoresis of polymer particles and bubbles. Microfluid. Nanofluid. 2012;13:625–635. doi: 10.1007/s10404-012-0979-6. [CrossRef] [Google Scholar]
  27. Gómez-Pastora J., González-Fernández C., Real E., Iles A., Bringas E., Furlani E.P., Ortiz I. Computational modeling and fluorescence microscopy characterization of a two-phase magnetophoretic microsystem for continuous-flow blood detoxification. Lab Chip. 2018;18:1593–1606. doi: 10.1039/C8LC00396C. [PubMed] [CrossRef] [Google Scholar]
  28. Forbes T.P., Forry S.P. Microfluidic magnetophoretic separations of immunomagnetically labeled rare mammalian cells. Lab Chip. 2012;12:1471–1479. doi: 10.1039/c2lc40113d. [PubMed] [CrossRef] [Google Scholar]
  29. Nandy K., Chaudhuri S., Ganguly R., Puri I.K. Analytical model for the magnetophoretic capture of magnetic microspheres in microfluidic devices. J. Magn. Magn. Mater. 2008;320:1398–1405. doi: 10.1016/j.jmmm.2007.11.024. [CrossRef] [Google Scholar]
  30. Plouffe B.D., Lewis L.H., Murthy S.K. Computational design optimization for microfluidic magnetophoresis. Biomicrofluidics. 2011;5:013413. doi: 10.1063/1.3553239. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  31. Hale C., Darabi J. Magnetophoretic-based microfluidic device for DNA isolation. Biomicrofluidics. 2014;8:044118. doi: 10.1063/1.4893772. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  32. Becker H., Gärtner C. Polymer microfabrication methods for microfluidic analytical applications. Electrophoresis. 2000;21:12–26. doi: 10.1002/(SICI)1522-2683(20000101)21:1<12::AID-ELPS12>3.0.CO;2-7. [PubMed] [CrossRef] [Google Scholar]
  33. Pekas N., Zhang Q., Nannini M., Juncker D. Wet-etching of structures with straight facets and adjustable taper into glass substrates. Lab Chip. 2010;10:494–498. doi: 10.1039/B912770D. [PubMed] [CrossRef] [Google Scholar]
  34. Wang T., Chen J., Zhou T., Song L. Fabricating microstructures on glass for microfluidic chips by glass molding process. Micromachines. 2018;9:269. doi: 10.3390/mi9060269. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  35. Castaño-Álvarez M., Pozo Ayuso D.F., García Granda M., Fernández-Abedul M.T., Rodríguez García J., Costa-García A. Critical points in the fabrication of microfluidic devices on glass substrates. Sens. Actuators B Chem. 2008;130:436–448. doi: 10.1016/j.snb.2007.09.043. [CrossRef] [Google Scholar]
  36. Prakash S., Kumar S. Fabrication of microchannels: A review. Proc. Inst. Mech. Eng. Part B J. Eng. Manuf. 2015;229:1273–1288. doi: 10.1177/0954405414535581. [CrossRef] [Google Scholar]
  37. Leester-Schädel M., Lorenz T., Jürgens F., Ritcher C. Fabrication of Microfluidic Devices. In: Dietzel A., editor. Microsystems for Pharmatechnology: Manipulation of Fluids, Particles, Droplets, and Cells. Springer; Basel, Switzerland: 2016. pp. 23–57. [Google Scholar]
  38. Bartlett N.W., Wood R.J. Comparative analysis of fabrication methods for achieving rounded microchannels in PDMS. J. Micromech. Microeng. 2016;26:115013. doi: 10.1088/0960-1317/26/11/115013. [CrossRef] [Google Scholar]
  39. Ng P.F., Lee K.I., Yang M., Fei B. Fabrication of 3D PDMS microchannels of adjustable cross-sections via versatile gel templates. Polymers. 2019;11:64. doi: 10.3390/polym11010064. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  40. Furlani E.P., Sahoo Y., Ng K.C., Wortman J.C., Monk T.E. A model for predicting magnetic particle capture in a microfluidic bioseparator. Biomed. Microdevices. 2007;9:451–463. doi: 10.1007/s10544-007-9050-x. [PubMed] [CrossRef] [Google Scholar]
  41. Tarn M.D., Peyman S.A., Robert D., Iles A., Wilhelm C., Pamme N. The importance of particle type selection and temperature control for on-chip free-flow magnetophoresis. J. Magn. Magn. Mater. 2009;321:4115–4122. doi: 10.1016/j.jmmm.2009.08.016. [CrossRef] [Google Scholar]
  42. Furlani E.P. Permanent Magnet and Electromechanical Devices; Materials, Analysis and Applications. Academic Press; Waltham, MA, USA: 2001. [Google Scholar]
  43. White F.M. Viscous Fluid Flow. McGraw-Hill; New York, NY, USA: 1974. [Google Scholar]
  44. Mathew B., Alazzam A., El-Khasawneh B., Maalouf M., Destgeer G., Sung H.J. Model for tracing the path of microparticles in continuous flow microfluidic devices for 2D focusing via standing acoustic waves. Sep. Purif. Technol. 2015;153:99–107. doi: 10.1016/j.seppur.2015.08.026. [CrossRef] [Google Scholar]
  45. Furlani E.J., Furlani E.P. A model for predicting magnetic targeting of multifunctional particles in the microvasculature. J. Magn. Magn. Mater. 2007;312:187–193. doi: 10.1016/j.jmmm.2006.09.026. [CrossRef] [Google Scholar]
  46. Furlani E.P., Ng K.C. Analytical model of magnetic nanoparticle transport and capture in the microvasculature. Phys. Rev. E. 2006;73:061919. doi: 10.1103/PhysRevE.73.061919. [PubMed] [CrossRef] [Google Scholar]
  47. Eibl R., Eibl D., Pörtner R., Catapano G., Czermak P. Cell and Tissue Reaction Engineering. Springer; Berlin/Heidelberg, Germany: 2009. [Google Scholar]
  48. Pamme N., Eijkel J.C.T., Manz A. On-chip free-flow magnetophoresis: Separation and detection of mixtures of magnetic particles in continuous flow. J. Magn. Magn. Mater. 2006;307:237–244. doi: 10.1016/j.jmmm.2006.04.008. [CrossRef] [Google Scholar]
  49. Alorabi A.Q., Tarn M.D., Gómez-Pastora J., Bringas E., Ortiz I., Paunov V.N., Pamme N. On-chip polyelectrolyte coating onto magnetic droplets-Towards continuous flow assembly of drug delivery capsules. Lab Chip. 2017;17:3785–3795. doi: 10.1039/C7LC00918F. [PubMed] [CrossRef] [Google Scholar]
  50. Zhang H., Guo H., Chen Z., Zhang G., Li Z. Application of PECVD SiC in glass micromachining. J. Micromech. Microeng. 2007;17:775–780. doi: 10.1088/0960-1317/17/4/014. [CrossRef] [Google Scholar]
  51. Mourzina Y., Steffen A., Offenhäusser A. The evaporated metal masks for chemical glass etching for BioMEMS. Microsyst. Technol. 2005;11:135–140. doi: 10.1007/s00542-004-0430-3. [CrossRef] [Google Scholar]
  52. Mata A., Fleischman A.J., Roy S. Fabrication of multi-layer SU-8 microstructures. J. Micromech. Microeng. 2006;16:276–284. doi: 10.1088/0960-1317/16/2/012. [CrossRef] [Google Scholar]
  53. Su N. 8 2000 Negative Tone Photoresist Formulations 2002–2025. MicroChem Corporation; Newton, MA, USA: 2002. [Google Scholar]
  54. Su N. 8 2000 Negative Tone Photoresist Formulations 2035–2100. MicroChem Corporation; Newton, MA, USA: 2002. [Google Scholar]
  55. Fu C., Hung C., Huang H. A novel and simple fabrication method of embedded SU-8 micro channels by direct UV lithography. J. Phys. Conf. Ser. 2006;34:330–335. doi: 10.1088/1742-6596/34/1/054. [CrossRef] [Google Scholar]
  56. Kazoe Y., Yamashiro I., Mawatari K., Kitamori T. High-pressure acceleration of nanoliter droplets in the gas phase in a microchannel. Micromachines. 2016;7:142. doi: 10.3390/mi7080142. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  57. Sharp K.V., Adrian R.J., Santiago J.G., Molho J.I. Liquid flows in microchannels. In: Gad-el-Hak M., editor. MEMS: Introduction and Fundamentals. CRC Press; Boca Raton, FL, USA: 2006. pp. 10-1–10-46. [Google Scholar]
  58. Oh K.W., Lee K., Ahn B., Furlani E.P. Design of pressure-driven microfluidic networks using electric circuit analogy. Lab Chip. 2012;12:515–545. doi: 10.1039/C2LC20799K. [PubMed] [CrossRef] [Google Scholar]
  59. Bruus H. Theoretical Microfluidics. Oxford University Press; New York, NY, USA: 2008. [Google Scholar]
  60. Beebe D.J., Mensing G.A., Walker G.M. Physics and applications of microfluidics in biology. Annu. Rev. Biomed. Eng. 2002;4:261–286. doi: 10.1146/annurev.bioeng.4.112601.125916. [PubMed] [CrossRef] [Google Scholar]
  61. Yalikun Y., Tanaka Y. Large-scale integration of all-glass valves on a microfluidic device. Micromachines. 2016;7:83. doi: 10.3390/mi7050083. [PMC free article] [PubMed] [CrossRef] [Google Scholar]
  62. Van Heeren H., Verhoeven D., Atkins T., Tzannis A., Becker H., Beusink W., Chen P. [(accessed on 9 March 2020)];Design Guideline for Microfluidic Device and Component Interfaces (Part 2) Version 3. Available online: http://www.makefluidics.com/en/design-guideline?id=7.
  63. Scheuble N., Iles A., Wootton R.C.R., Windhab E.J., Fischer P., Elvira K.S. Microfluidic technique for the simultaneous quantification of emulsion instabilities and lipid digestion kinetics. Anal. Chem. 2017;89:9116–9123. doi: 10.1021/acs.analchem.7b01853. [PubMed] [CrossRef] [Google Scholar]
  64. Lynch E.C. Red blood cell damage by shear stress. Biophys. J. 1972;12:257–273. [PMC free article] [PubMed] [Google Scholar]
  65. Paul R., Apel J., Klaus S., Schügner F., Schwindke P., Reul H. Shear stress related blood damage in laminar Couette flow. Artif. Organs. 2003;27:517–529. doi: 10.1046/j.1525-1594.2003.07103.x. [PubMed] [CrossRef] [Google Scholar]
  66. Gómez-Pastora J., Karampelas I.H., Xue X., Bringas E., Furlani E.P., Ortiz I. Magnetic bead separation from flowing blood in a two-phase continuous-flow magnetophoretic microdevice: Theoretical analysis through computational fluid dynamics simulation. J. Phys. Chem. C. 2017;121:7466–7477. doi: 10.1021/acs.jpcc.6b12835. [CrossRef] [Google Scholar]
  67. Lim J., Yeap S.P., Leow C.H., Toh P.Y., Low S.C. Magnetophoresis of iron oxide nanoparticles at low field gradient: The role of shape anisotropy. J. Colloid Interface Sci. 2014;421:170–177. doi: 10.1016/j.jcis.2014.01.044. [PubMed] [CrossRef] [Google Scholar]
  68. Culbertson C.T., Sibbitts J., Sellens K., Jia S. Fabrication of Glass Microfluidic Devices. In: Dutta D., editor. Microfluidic Electrophoresis: Methods and Protocols. Humana Press; New York, NY, USA: 2019. pp. 1–12. [Google Scholar]
Figure 20. Top: image of electrospray, bottom: cone-jet profile using the CF emitter. Distance between the carbon fiber tip and the counter electrode is 4.0 mm, potential difference is 3500 V, flow rate is 300 nL min−1 .

Modeling and characterization of a carbon fiber emitter for electrospray ionization

A K Sen1, J Darabi1, D R Knapp2 and J Liu2
1 MEMS and Microsystems Laboratory, Department of Mechanical Engineering,
University of South Carolina, 300 Main Street, Columbia, SC 29208, USA
2 Department of Pharmacology, Medical University of South Carolina, 173 Ashley Avenue,
Charleston, SC 29425, USA
E-mail: darabi@engr.sc.edu

뾰족한 탄소 섬유(CF)를 사용하는 새로운 마이크로 스케일 이미터는 질량 분석 (MS) 분석에서 전기 분무에 사용할 수 있습니다. 탄소 섬유는 360 µm OD 및 75 µm ID의 용융 실리카 모세관과 동축에 위치하며 날카로운 팁은 튜브 말단에서 30 µm 연장됩니다.

Abstract

전기 분무 이온화 (ESI) 프로세스는 전기 유체 역학을 해결하기 위한 Taylor–Melcher 누설 유전체 유체 모델 및 액체-가스 인터페이스 추적을 위한 유체 부피 (VOF) 접근 방식을 기반으로 하는 전산 유체 역학 (CFD) 코드를 사용하여 시뮬레이션 됩니다. CFD 코드는 먼저 기존 지오메트리에 대해 검증한 다음 CF 이미터 기반 ESI 모델을 시뮬레이션하는데 사용됩니다.

시뮬레이션된 전류 흐름 및 전류 전압 결과는 CF 이미터의 실험 결과와 잘 일치합니다. 이미터 형상, 전위차, 유속 및 액체의 물리적 특성이 CF 이미터의 전기 분무 거동에 미치는 영향을 철저히 조사합니다.

스프레이 전류와 제트 직경은 액체의 유속, 전위차 및 물리적 특성과 상관 관계가 있으며 상관 결과는 문헌에 보고된 결과와 정량적으로 비교됩니다. (이 기사의 일부 그림은 전자 버전에서만 색상입니다)

Introduction

1980 년대 후반부터 매트릭스 보조 레이저 탈착 이온화 (MALDI)와 전기 분무 이온화 (ESI)의 두 가지 이온화 기술을 구현하여 감도, 속도 및 구조 정보 수준 측면에서 MS 분석이 엄청나게 성장했습니다. 1980 년대 초까지 전자 충격 (EI) 또는 화학 이온화 (CI) 방법은 가스 크로마토 그래피에 적합한 작은 생체 분자를 이온화 하는 데 사용되었습니다.

그러나 크고 열에 민감한 비 휘발성 샘플은 적절한 사전 처리 없이 EI 또는 CI-MS 기술로 분석 할 수 없습니다 [1]. ESI 기술을 사용하면 액체상에서 직접 이러한 큰 분자를 분석 할 수 있습니다 [2]. Zeleny [3, 4]는 출구에 높은 전위를 적용하여 모세관에서 액체 용액을 분사 할 수 있음을 보여주었습니다.

Dole [5, 6] 및 Fenn [7]의 선구적인 연구는 ESI를 고분자 및 생체 분자와 같은 대형 화합물의 이온화 방법으로 표시했습니다. 이에 이어이 기술에 의한 기상 이온 발생에 관련된 과정과 메커니즘이 널리 조사되고 있습니다.

ESI 방법에서 기체 이온화 된 분자는 강한 전계가 있는 상태에서 미세한 물방울을 생성하여 액체 용액에서 생성됩니다. ESI 프로세스의 이러한 능력은 단백질 및 기타 생체 분자 연구에 자연적으로 적용됨을 발견했습니다. ESI 방법과 관련된 다양한 프로세스가 그림 1에 나와 있습니다.

Figure 1. Schematic of an ESI process.
Figure 1. Schematic of an ESI process.

ESI 전위는 일반적으로 전도성 물질로 코팅 된 이미 터 튜브를 통해 외부에서 샘플 액체에 적용되지만 액체 샘플 내부에 적용될 수도 있습니다. Herring과 Qin [8]은 이미 터 팁에 삽입된 팔라듐 와이어를 통해 전기 분무 전위가 적용되는 모세관 전기 영동 (CE)을위한 ESI 인터페이스를 보여주었습니다.

Chiou의 설계 [9]에서는 작은 PDMS 칩에 있는 샘플 저장소, 마이크로 채널 및 실리카 모세관 노즐과 통합 된 내장 전극을 통해 전기 분무를 위한 고전압이 적용되었습니다.

Cao and Moini [10]는 ESI 전압이 모세관 내부에 위치한 전극을 통해인가되고 전기적 접촉이 출구 근처 모세관 벽의 작은 구멍을 통해 유지되는 전기 분무 방출기를 설계했습니다. 작은 모세관 직경 (~ 10 µm)을 가진 이미 터를 사용하여 낮은 전압에서 전기 분무가 가능하지만, 더 작은 구멍은 과도한 배압으로 인해 쉽게 막힐 수 있습니다.

직경이 더 큰 (> 50µm) 이미 터를 처리하는 것이 더 쉽습니다. 그러나 그들은 더 작은 직경의 이미 터만큼 효율적이지 않습니다 [11]. 일반적으로 ESI 전압을 적용하기 위해 유리 또는 용융 실리카와 같은 절연 재료로 제작 된 저 유량 이미 터의 외주에 전도성 코팅이 적용됩니다.

용융 실리카 모세관의 끝 부분에있는 스퍼터 코팅 된 귀금속 층은 내구성에 빠르게 영향을 미치는 것으로 관찰되었습니다. 코팅의 빠른 열화는 방전, 전기 화학적 반응 및 층과 용융 실리카 표면 사이의 불량한 기계적 결합으로 인해 발생할 수 있습니다.

이러한 에미 터의 수명은 스퍼터 코팅 후에 금을 전기 도금하거나 [12] 스퍼터 코팅 된 금 위에 SiOx를 코팅하여 증가시킬 수 있습니다 [13]. 크롬 또는 니켈 합금의 접착층 위에 금으로 코팅 된 이미 터는 우수한 결합력을 제공 할 수 있으며 음극으로 작동 할 때 내구성이 있습니다.

그러나 양극으로 작동하는 동안 접착층은 금 막을 통해 화학적으로 용해됩니다. 이미 터의 안정성과 내구성을 향상시키기 위해 대체 전도성 코팅이 평가되었습니다.

안정적인 ESI 작동을 위해 콜로이드 흑연 코팅 이미 터가 사용되었으며 수명이 길었습니다 [14]. 폴리아닐린 (PANI) 코팅 이미 터는 두꺼운 코팅으로 인해 높은 내구성을 보여주고 방전에 강합니다. PANIcoated와 gold-coated nanospray emitter의 electrospray ionization 거동을 비교 한 결과 PANIcoated emitter는 goldcoated emitter와 비슷한 향상된 감도를 제공합니다 [15].

그라파이트-폴리이 미드 혼합물은 또한 무 접착 전기 분무 방출기의 경우 전도성 코팅으로 사용되었습니다. 전도성 코팅의 안정성은 산화 스트레스 동안 좋은 성능을 나타내는 전기 화학적 방법에 의해 조사되었습니다 [16].

탄소 코팅 이미 터의 기능은 마이크로 스프레이 및 시스리스 CE 및 ESI 응용 분야에서 입증되었습니다. 이 이미 터는 견고하지는 않지만 방수가 되지 않는 CE 또는 ESI 애플리케이션에 충분히 내구성이있었습니다 [17].

우리는 막힘 문제를 제거하고 시료 액체와 금층 사이의 접촉 문제를 피할 수있는 뾰족한 탄소 섬유 기반의 새로운 ESI 방출기를 도입하여 ESI 시스템의 적용 성, 신뢰성 및 내구성을 향상 시켰습니다 [18]. 이 작업에서 탄소 섬유 기반 ESI 이미 터는 전산 유체 역학 (CFD) 소프트웨어 패키지 FLOW-3D [19]를 사용하여 시뮬레이션됩니다.

실험은 새로운 CF 이미 터를 사용하여 수행됩니다. 모델 예측은 실험 결과와 비교됩니다. 새로운 이미 터의 ESI 성능은 이미 터의 기하학적 구조, 유속, 액체의 물리적 특성과 같은 다양한 매개 변수에 대한 반응을 연구하여 평가됩니다.

스프레이 전류 및 제트 직경은 유량 및 액체의 특성과 상관 관계가 있으며 상관 결과는 문헌에보고 된 결과와 정량적으로 비교됩니다. 다음 섹션에서 ESI 공정을 지배하는 전기 유체 역학 이론은 Taylor–Melcher 누설 유전체 모델 [20]을 참조하여 설명됩니다.

그런 다음 Hartman 등이 사용하는 ESI 구성을 고려하여 CFD 코드의 유효성을 확인합니다 [21]. 또한 CF 기반 ESI 모델에 대한 시뮬레이션 및 실험 결과가 제시되고 논의됩니다. 마지막으로 모수 연구 결과와 상관 관계를 제시하고 논의합니다.

Figure 2. Forces in the liquid cone.
Figure 2. Forces in the liquid cone.
Figure 3. Schematic of the ESI model studied by Hartman et al [21].
Figure 3. Schematic of the ESI model studied by Hartman et al [21].
Figure 6. Cone-Jet profile and the electric potential contours at 19 kV; cone length is 4.3 mm.
Figure 6. Cone-Jet profile and the electric potential contours at 19 kV; cone length is 4.3 mm.
Figure 7. A photograph of the experimental cone shape; cone length is 4.2 ± 0.2 mm [21].
Figure 7. A photograph of the experimental cone shape; cone length is 4.2 ± 0.2 mm [21].
Figure 15. Electric field contours at various time steps
Figure 15. Electric field contours at various time steps
Figure 20. Top: image of electrospray, bottom: cone-jet profile using the CF emitter. Distance between the carbon fiber tip and the counter electrode is 4.0 mm, potential difference is 3500 V, flow rate is 300 nL min−1 .
Figure 20. Top: image of electrospray, bottom: cone-jet profile using the CF emitter. Distance between the carbon fiber tip and the counter electrode is 4.0 mm, potential difference is 3500 V, flow rate is 300 nL min−1 .

References

[1] Siuzdak M 1996 Mass Spectrometry for Biotechnology (New York: Academic)
[2] Cole R B (ed) 1997 Electrospray Ionization Mass Spectrometry (New York: Wiley-Interscience)
[3] Zeleny J 1914 Phys. Rev. 3 69–91
[4] Zeleny J 1917 Phys. Rev. 10 1–6
[5] Dole M, Mack L L, Hines R L, Mobley R C, Ferguson L D and Alice M B 1968 Molecular beams of macroions
J. Chem. Phys. 49 2240–9
[6] Clegg G A and Dole M 1971 Molecular beams of macroions: III. Zein and polyvinylpyrrolidone Biopolymers
10 821–6
[7] Fenn J B, Mann M, Meng C K, Wong S F and Whitehouse C M 1989 Electrospray ionization for mass
spectrometry of large biomolecules Science 246 64–71
[8] Herring C J and Qin J 1999 An on-line preconcentrator and the evaluation of electrospray interfaces for the capillary
electrophoresis/mass spectrometry of peptides Rapid Commun. Mass Spectr. 13 1–7
[9] Chiou C H, Lee G B, Hsu H T, Chen P W and Liao P C B 2002 Microscale Tools for Sample Preparation, Separation
and Detection of Neuropeptides Sensors Actuators B 86 280–6
[10] Cao P and Moini M 1997 A novel sheathless interface for capillary electrophoresis/electrospray ionization mass
spectrometry using an in-capillary electrode J. Am. Soc. Mass Spectrom 8 561–4
[11] Janini G M, Conards T P, Wilkens K L, Issaq H J and Veenstra T D 2003 A sheathless nanoflow electrospray
interface for on-line capillary electrophoresis mass spectrometry Anal. Chem 75 1615–9
[12] Barroso M B de Jong and Ad P 1999 Sheathless preconcentration-capillary zone electrophoresis-mass
spectrometry applied to peptide analysis J. Am. Soc. Mass Spectrom 10 1271–8
[13] Valaskovic G A and McLafferty F W 1996 Long-lived metallized tips for nanoliter electrospray mass spectrometry
J. Am. Soc. Mass Spectrom. 7 1270–2
[14] Zhu X, Thiam S, Valle B C and Warner I M 2002 A colloidal graphite coated emitter for seathless capillary
electrophoresis/nanoelectrospray ionization mass spectrometry Anal. Chem 74 5405–9
[15] Maziarz E P I II, Lorenz S A, White T P and Wood T D 2000 Polyaniline: a conductive polymer coating for durable
nanospray emitters J. Am. Soc. Mass. Spectrom 11 659–63
[16] Nilsson S, Wetterhall M, Bergquist J, Nyholm L and Markides K E 2001 A simple and robust conductive
graphite coating for sheathless electrospray emitters used in capillary electrophoresis/mass spectrometry Rapid
Commun. Mass Spectr. 15 1997–2000
[17] Chang Y Z and Her G R 2000 Sheathless capillary electrophoresis/electospray mass spectrometry using a
carbon-coated tapered fused silica capillary with a beveled edge Anal. Chem. 72 626–30
[18] Liu J, Ro K W, Busman M and Knapp D R 2004 Electrospray ionization with a pointed carbon fiber emitter Anal. Chem. 76 3599–606
[19] Hirt C W 2004 Electro-hydrodynamics of semi–conductive fluids: with application to electro–spraying Flow Science
Technical Note 70 FSI–04–TN70 1–7
[20] Saville D A 1997 Electrohydrodynamcis: the Taylor–Melcher leaky dielectric model Annu. Rev. Fluid Mech. 29 27–64
[21] Hartman R P A, Brunner D J, Camelot D M A, Marijnissen J C M and Scarlett B 1999
Electrohydrodynamic atomization in the cone-jet mode physical modeling of the liquid cone and jet J. Aerosol Sci.
30 823–49
[22] Castellanos A 1998 Basic Concepts and Equations in Electrohydrodynamics Electrohydrodynamics
ed A Castellanos (Berlin: Springer)
[23] Melcher J R 1981 Continuum Electromechanics (Cambridge, MA: MIT Press)
[24] Hirt C W and Nichols B D 1981 Volume of fluid (VOF) method for the dynamics of free boundaries J. Comp. Phys.
39 201–25
[25] De la Mora F J and Loscertales I G 1994 The current emitted by highly conducting Taylor cones J. Fluid Mech. 260
155–84
[26] Ganan-Calvo A M 1997 Cone–jet analytical extension of Taylor’s electrostatic solution and the asymptotic universal
scaling laws in electrospraying Phys. Rev. Lett. 79 217–20
[27] Higuera F J 2004 Current/flow–rate characteristic of an electrospray with a small meniscus J. Fluid Mech.
513 239–46
[28] Zeng J, Sobek D and Korsmeyer T Electro-hydrodynamic modeling of electrospray ionization: cad for a microfluidic
device-mass spectrometer interface Transducers ’03: 12th Int. Conf. on Solid State Sensors, Actuators and
Microsystems 2 1275–8
[29] Ganan–Calvo A M, Davila J and Barrero A 1997 Current and droplet size in the electrospraying of liquids. Scaling laws J. Aerosol Sci. 28 249–75
[30] Cloupeau M and Prunet-Foch B 1989 Electrostatic spraying of liquids in cone–jet mode J. Electrost. 22 135–59

Fig. 7. Simulation results of temperature distribution between Ni stamps and PBO-SAM/PMMA substrate in NIL process: (A) stamp cross-sectional, (B) PMMA substrate cross-sectional, (C) 3-dimensional and (D) intrinsic 3-dimensional views, respectively. The study of computed condition in nanoimprint process is at 150 o C and 50 bar during 10 min. Note that for NIL experimental parameters, the simulated results have already decided before doing nanoimprint experiment.

A non-fluorine mold release agent for Ni stamp in nanoimprint process

Tien-Li Chang a,*, Jung-Chang Wang b
, Chun-Chi Chen c
, Ya-Wei Lee d
, Ta-Hsin Chou a
a Mechanical and Systems Research Laboratories, Industrial Technology Research Institute, Rm. 125, Building 22, 195 Section 4, Chung Hsing Road, Chutung, Hsinchu 310, Taiwan, ROC bDepartment of Manufacturing Research and Development, ADDA Corporation, Taiwan
cNational Nano Device Laboratories, Taiwan
d Research and Development Division, Ordnance Readiness Development Center, Taiwan

Abstract

이 연구는 나노 임프린트 공정에서 Ni 몰드 스탬프와 PMMA (폴리 메틸 메타 크릴 레이트) 기판 사이의 접착 방지 층으로서 새로운 재료를 제시합니다. 폴리 벤족 사진 ((6,6′-bis (2,3-dihydro3-methyl-4H-1,3-benzoxazinyl))) 분자 자기 조립 단층 (PBO-SAM)은 점착 방지 코팅제로 간주되어 불소 함유 화합물은 Ni / PMMA 기판의 나노 임프린트 공정을 개선 할 수 있습니다. 이 작업에서 나노 구조 기반 Ni 스탬프와 각인 된 PMMA 몰드는 각각 전자빔 석판화 (EBL)와 수제 나노 임프린트 장비에 의해 수행됩니다. 제작 된 나노 패턴의 형성을 제어하기 위해 시뮬레이션은 HEL (hot embossing lithography) 공정 동안 PBO-SAM / PMMA 기판의 변형에 대한 온도 분포의 영향을 분석 할 수 있습니다. 여기서 기둥 패턴의 직경은 Ni 스탬프 표면에 200nm 및 400nm 피치입니다. 이 적합성 조건에서 소수성 PBO-SAM 표면을 기반으로하여 Ni 몰드 스탬프의 결과는 품질 및 수량 제어에서 90 % 이상의 개선을 추론합니다.

Introduction

나노 임프린트 리소그래피 (NIL)는 초 미세 패터닝 기판 기술을 대량 생산할 수있는 가장 큰 잠재력입니다 [1,2]. 최근에는 광전자 장치 [3], 양자 컴퓨팅 장치 [4], 바이오 센서 [5] 및 전자 장치 [6]에 요구 될 수있는 NEMS / MEMS 기술의 빠른 개발이 이루어지고 있습니다.

따라서 기존의 포토 리소 그래프는 할당에 적합한 방법이 아닐 수 있습니다 [7]. X 선, 이온빔, 전자빔 리소그래피의 경우 LCD의 도광판 초박막 판과 같은 대 면적 패턴 제작에 적합하지 않습니다. 제어하기 어렵습니다. 일부 제작된 문제를 기반으로 NIL 프로세스는 재료, 패턴 크기, 구조 및 기판 지형면에서 유연성을 제공합니다 [8].

오늘날 NIL 제조 방법은 낮은 비용과 높은 처리량의 높은 패터닝 해상도의 조합으로 학제 간 나노 스케일 연구 및 상용 제품의 새로운 문을 열 수 있는 큰 관심을 받고 있습니다. 그러나 이 나노 임프린트 기술이 산업 규모 공정을 위해 충분히 성숙하기 전에 몇 가지 응용 문제를 해결해야 합니다.

각인된 몰드 공정은 종종 고온 (폴리머의 유리 전이 온도에 대해> 100oC)과 고압 (> 100bar)에서 수행되기 때문에 분명히 바람직하지 않습니다. 가열 및 냉각 공정의 열주기는 금형 및 각인 된 기판의 왜곡을 유발할 수 있습니다. 한 가지 특별한 문제는 스탬프와 폴리머 사이의 접착 방지 층 처리를 제어하여 기계적 결함이 임프린트 품질과 스탬프 수명에 영향을 미칠 수있는 중요한 패턴 결함이되는 것을 방지하는 것입니다.

Schift et al. 플루오르화 트리클로로 실란을 마이크로 미터 체제에서 실리콘에 대한 접착 방지 코팅으로 사용하는 것으로 입증되었습니다 [9]. 또한 Park et al. Ni 몰드 스탬프에 더 나은 접착 방지 코팅 공정을 달성하기 위해 불소화 실란제를 사용했습니다 [10].

그러나 지금까지 Ni 스탬프에 대한 접착 방지 코팅 처리의 NIL 공정에서 비 불소 물질에 대한 시도는 거의 이루어지지 않았습니다. 우리의 생활 환경은 그것을 유지하기 위해 불소가 아닌 물질이 필요합니다. 또한 Ni 계 소재의 부드러운 특성을 바탕으로 가장 중요한 롤러 나노 임프린트 기술을 개발할 수 있습니다.

본 연구의 목적은 Ni 스탬프와 PMMA 기판 사이의 점착 방지 코팅제로 PBO-SAM을 개발하여 나노 제조 기술, 즉 NIL을 향상시키는 것입니다.

Experiment

먼저 4,4′- 이소 프로필 리 덴디 페놀 (비스페놀 -A, BA-m), 포름 알데히드 및 ​​메틸 아민을 반응시켜 폴리 벤족 사진을 제조 하였다. 미국 Aldrich Chemical company, Inc.에서 구입 한 모든 화학 물질. 합성 과정에서 포름 알데히드/디 옥산 및 메틸 아민 / 디 옥산 물질을 10 o C에서 항아리에서 10분 동안 측정하는 벤족 사진 단량체가 필요했습니다.

디 에틸 에테르를 기화시킨 후, 벤족 사진 전구체가 완성되었다. benzoxazine 전구체를 140 o C에서 1 시간 동안 가열하면 BA-m 폴리 벤족 사진을 얻을 수 있습니다. 다음으로 4 인치입니다.

이 연구에서는 p 형 Si (10 0) 웨이퍼를 사용할 수 있습니다. SiO2 기반 Ni (원자량 5.87g / mole) 기판의 제조를 위해 Ti (5nm) 및 SiO2 (20nm)를 순차적으로 증착 한 후 O2- 플라즈마 처리를 수행했습니다. Ni 기판과 SiO2 층 사이의 접착력을 높이기 위해 Ti 중간층이 사용되었습니다. 아세톤, 이소프로판올 및 탈 이온수를 사용하여 세척 한 후 샘플을 포토 레지스트 (ZEP520A-7, Nippon Zeon Co., Ltd.)로 스핀 코팅했습니다.

Fig. 1. Schematic diagram of nanostructures using NIL process: (A) EBL equipment for fabricated mold stamp. (B) HEL equipment for nanoimprint pattern with computer controlled electronics. (C) A nickel-based pillar mold can imprint into a PBO-SAM polymer resist layer; afterward, the mold removal and pattern transfer are based on anisotropic etching to remove reside.
Fig. 1. Schematic diagram of nanostructures using NIL process: (A) EBL equipment for fabricated mold stamp. (B) HEL equipment for nanoimprint pattern with computer controlled electronics. (C) A nickel-based pillar mold can imprint into a PBO-SAM polymer resist layer; afterward, the mold removal and pattern transfer are based on anisotropic etching to remove reside.

마스터 몰드는 그림 1 (A)에서 Ni 필름의 반응성 이온 에칭 (RIE)과 함께 Crestec CABL8210 전자 빔 직접 쓰기 도구 (30 keV, 100 pA)를 사용하여 제작되었습니다. 그런 다음 시뮬레이션된 결과는 NIL 프로세스에서 엠보싱 압력으로 기계적 고장의 효과를 제공할 수 있으며, 이는 우리가 원하는 나노 패턴 설계 및 연구에 도움이 될 수 있습니다.

PBOSAM / PMMA 기판 모델의 변형은 3 차원 접근법에 기반한 유한 체적 방법 (FVM)을 통해 예측할 수 있습니다. Navier-Stokes 방정식 [11]에서 압력과 속도 사이의 결합은 SIMPLE 알고리즘을 사용하여 이루어집니다. 2 차 상향 이산화 방식은 대류 플럭스 및 운동량의 확산 플럭스, 유체의 질량 분율에 대한 중심 차이 방식에 대해 구현됩니다. 완화 부족 요인의 일반적인 값은 0.5입니다.

수렴 기준이 1105로 설정된 연속성을 제외한 모든 변수에 대해 잔차가 1103 미만인 경우 솔루션이 수렴된 것으로 간주됩니다. 여기서 각인된 나노 패턴은 그림 1 (B)와 같이 수제 장비에서 수행한 HEL 공정을 통해 사용할 수 있습니다. PBO-SAM 코팅 방법으로 HEL 절차를 활용 한 나노 패턴의 제작은 그림 1 (C)에 개략적으로 표시되었습니다.

200nm의 얇은 PMMA 필름 (분자량 15kg / mole)을 SiO2 기판에 스핀 코팅 한 후 160oC에서 30 분 동안 핫 플레이트에서 베이킹했습니다. 또한 PBO-SAM 코팅은 접착 방지제입니다. CVD 공정에 의해 증착되었습니다. 마스터는 150oC 및 50bar에서 10 분 동안 PBO-SAM / PMMA 기판 필름에 엠보싱하여 복제되었습니다.

마지막으로, 엠보싱 된 나노 구조물의 바닥에 남아 있던 PBO-SAM / PMMA 층은 RIE 처리로 제거되었습니다. 각 임프린트 후 스탬프 및 기판의 품질이 제작 된 후 현미경을 사용하여 관찰하고 물 접촉각 (CA) 측정을 사용하여 습윤 및 접착 특성을 알아낼 수 있습니다.

Fig. 2. FTIR absorption spectrum of polybenzoxazines indicates the vibrational modes of molecular bonds.
Fig. 2. FTIR absorption spectrum of polybenzoxazines indicates the vibrational modes of molecular bonds.
Fig. 3. FE-SEM micrograph of Ni stamps before imprinted PMMA substrate. The pillar diameter is 200 nm, and its period is 400 nm.
Fig. 3. FE-SEM micrograph of Ni stamps before imprinted PMMA substrate. The pillar diameter is 200 nm, and its period is 400 nm.
Fig. 5. Contact angles of water drops on (A) a PMMA polymer film surface, and (B) a smooth PBO-SAM coating film surfaceFig. 6. Simulation of Ni stamps and PBO-SAM/PMMA substrate in NIL process: (A) A nanoimprint system geometry, and (B) its grid plot.
Fig. 5. Contact angles of water drops on (A) a PMMA polymer film surface, and (B) a smooth PBO-SAM coating film surfaceFig. 6. Simulation of Ni stamps and PBO-SAM/PMMA substrate in NIL process: (A) A nanoimprint system geometry, and (B) its grid plot.
Fig. 7. Simulation results of temperature distribution between Ni stamps and PBO-SAM/PMMA substrate in NIL process: (A) stamp cross-sectional, (B) PMMA substrate cross-sectional, (C) 3-dimensional and (D) intrinsic 3-dimensional views, respectively. The study of computed condition in nanoimprint process is at 150 o C and 50 bar during 10 min. Note that for NIL experimental parameters, the simulated results have already decided before doing nanoimprint experiment.
Fig. 7. Simulation results of temperature distribution between Ni stamps and PBO-SAM/PMMA substrate in NIL process: (A) stamp cross-sectional, (B) PMMA substrate cross-sectional, (C) 3-dimensional and (D) intrinsic 3-dimensional views, respectively. The study of computed condition in nanoimprint process is at 150 o C and 50 bar during 10 min. Note that for NIL experimental parameters, the simulated results have already decided before doing nanoimprint experiment.

References

[1] M.D. Austin, H.X. Ge, W. Wu, M.T. Li, Z.N. Yu, D. Wasserman, S.A. Lyon, S.Y. Chou, Nature 417 (2002) 835.
[2] S.Y. Chou, C. Keimel, J. Gu, Appl. Phys. Lett. 84 (2004) 5299.
[3] Q. Wang, G. Farrell, P. Wang, G. Rajan, T. Thomas, Sensor Actuator A 134 (2007) 405.
[4] C. Kentsch, W. Henschel, D. Wharam, D.P. Kern, Microelectron. Eng. 83 (2006) 1753.
[5] T.L. Chang, Y.W. Lee, C.C. Chen, F.H. Ko, Microelectron. Eng. 84 (2007) 1689.
[6] S. Tisa, F. Zappa, A. Tosi, S. Cova, Sensor Actuator A 140 (2007) 113.
[7] M. Agirregabiria, F.J. Blanco, J. Berganzo, M.T. Arroyo, A. Fullaondo, K. Mayora, J.M. Ruano-López, Lab Chip 5 (2005) 5545.
[8] W. Hu, E.K.F. Yim, R.M. Reano, K.W. Leong, S.W. Pang, J. Vac. Sci. Technol. B 84 (2005) 2984.
[9] H. Schift, L.J. Heyderman, C. Padeste, J. Gobrecht, Microelectron. Eng. 423 (2002) 61.
[10] S. Park, H. Schift, C. Padeste, B. Schnyder, R. Kötz, J. Gobrecht, Microelectron. Eng. 73–74 (2004) 196.
[11] A. Yokoo, M. Nakao, H. Yoshikawa, H. Masuda, T. Tamamura, Jpn. J. Appl. Phys. 38 (1999) 7268.

Figure 2. Simulation of droplet separation by EWOD

Non-Linear Electrohydrodynamics in Microfluidic Devices

미세 유체 장치의 비선형 전기 유체 역학

by Jun ZengHewlett-Packard Laboratories, Hewlett-Packard Company, 1501 Page Mill Road, Palo Alto, CA 94304, USAInt. J. Mol. Sci.201112(3), 1633-1649; https://doi.org/10.3390/ijms12031633Received: 24 January 2011 / Revised: 10 February 2011 / Accepted: 24 February 2011 / Published: 3 March 2011

Abstract

Since the inception of microfluidics, the electric force has been exploited as one of the leading mechanisms for driving and controlling the movement of the operating fluid and the charged suspensions. Electric force has an intrinsic advantage in miniaturized devices. Because the electrodes are placed over a small distance, from sub-millimeter to a few microns, a very high electric field is easy to obtain. The electric force can be highly localized as its strength rapidly decays away from the peak. This makes the electric force an ideal candidate for precise spatial control. The geometry and placement of the electrodes can be used to design electric fields of varying distributions, which can be readily realized by Micro-Electro-Mechanical Systems (MEMS) fabrication methods. In this paper, we examine several electrically driven liquid handling operations. The emphasis is given to non-linear electrohydrodynamic effects. We discuss the theoretical treatment and related numerical methods. Modeling and simulations are used to unveil the associated electrohydrodynamic phenomena. The modeling based investigation is interwoven with examples of microfluidic devices to illustrate the applications. 

Keywords: dielectrophoresiselectrohydrodynamicselectrowettinglab-on-a-chipmicrofluidicsmodelingnumerical simulationreflective display

요약

미세 유체학이 시작된 이래로 전기력은 작동 유체와 충전 된 서스펜션의 움직임을 제어하고 제어하는 ​​주요 메커니즘 중 하나로 활용되어 왔습니다. 전기력은 소형 장치에서 본질적인 이점이 있습니다. 전극이 밀리미터 미만에서 수 미크론까지 작은 거리에 배치되기 때문에 매우 높은 전기장을 쉽게 얻을 수 있습니다. 

전기력은 강도가 피크에서 멀어지면서 빠르게 감소하기 때문에 고도로 국부화 될 수 있습니다. 이것은 전기력을 정밀한 공간 제어를 위한 이상적인 후보로 만듭니다.

전극의 기하학적 구조와 배치는 다양한 분포의 전기장을 설계하는 데 사용될 수 있으며, 이는 MEMS (Micro-Electro-Mechanical Systems) 제조 방법으로 쉽게 실현할 수 있습니다. 

이 논문에서 우리는 몇 가지 전기 구동 액체 처리 작업을 검토합니다. 비선형 전기 유체 역학적 효과에 중점을 둡니다. 이론적 처리 및 관련 수치 방법에 대해 논의합니다. 모델링과 시뮬레이션은 관련된 전기 유체 역학 현상을 밝히는 데 사용됩니다. 모델링 기반 조사는 응용 분야를 설명하기 위해 미세 유체 장치의 예와 결합됩니다. 

키워드 : 유전 영동 ; 전기 유체 역학 ; 전기 습윤 ; 랩 온어 칩 ; 미세 유체 ; 모델링 ; 수치 시뮬레이션 ; 반사 디스플레이

Droplet processing array Droplet based BioFlip
igure 1. Example of droplet-based digital microfluidics architecture. Above is an elevation view showing the layered structure of the chip. Below is a diagram illustrating the system (Adapted from [4]).
Figure 2. Simulation of droplet separation by EWOD
Figure 2. Simulation of droplet separation by EWOD. The top two figures illustrate the device configuration. Electric voltages are applied to all four electrodes embedded in the insulating material. The bottom left figure shows transient simulation solution. It illustrates the process of separating one droplet into two via EWOD. The bottom right figure shows the electric potential distribution inside the device. The color indicates the electric potential; the iso-potential surfaces are also drawn. The image shows the electric field is absent within the droplet body indicating the droplet is either conductive or highly polarizable.
Figure 4. Transient sequence of the Taylor cone formation
Figure 4. Transient sequence of the Taylor cone formation: simulation and experiment comparison. Experimental images are shown in the top row. Simulation results are shown in the bottom row. Their correspondence is indicated by the vertical alignment (Adapted from [4]).
Figure 6. Simulation of charge screening effect using a parallel-plate cell
Figure 6. Simulation of charge screening effect using a parallel-plate cell. Top-left image shows the electric current as function of time and driving voltage, top-right image shows the evolution of the species concentration as function of time and space, the bottom image shows the electric current readout after switching the applied voltage.
Figure 7. Transient simulation of electrohydrodynamic instability and the development of the cellular convective flow pattern.
Figure 7. Transient simulation of electrohydrodynamic instability and the development of the cellular convective flow pattern.
Figure 3. Simulation of dielectrophoresis driven axon migration
Figure 3. Simulation of dielectrophoresis driven axon migration. The set of small images on the left shows a transient simulation of single axon migration under an electric field generated by a pin electrode. The image on the right is a snapshot of a simulation where two axons are fused by dielectrophoresis using a pin electrode. Axons are outlined in white. Also shown are the iso-potential curves.

References

  1. Muller, RS. MEMS: Quo vadis in century XXI. Microelectron. Eng 200053(1–4), 47–54. [Google Scholar]
  2. Reyes, DR; Iossifidis, D; Auroux, PA; Manz, A. Micro total analysis systems. 1. Introduction, theory, and technology. Anal.Chem 200274, 2623–2636. [Google Scholar]
  3. Levy, U; Shamai, R. Tunable optofluidic devices. Microfluid. Nanofluid 20084, 97–105. [Google Scholar]
  4. Zeng, J; Korsmeyer, FT. Principles of droplet electrohydrodynamics for lab-on-a-chip. Lab Chip 20044, 265–277. [Google Scholar]
  5. Fair, RB. Digital microfluidics: Is a true lab-on-a-chip possible? Microfluid. Nanofluid 20073, 245–281. [Google Scholar]
  6. Pollack, MG; Fair, RB; Shenderov, AD. Electrowetting-based actuation of liquid droplets for microfluidic applications. Appl. Phys. Lett 200077(11), 1725–1726. [Google Scholar]
  7. Peykov, V; Quinn, A; Ralston, J. Electrowetting: A model for contact-angle saturation. Colloid Polym. Sci 2000278, 789–793. [Google Scholar]
  8. Verheijen, HJJ; Prins, MWJ. Reversible electrowetting and trapping of charge: Model and experiments. Langmuir 199915, 6616–6620. [Google Scholar]
  9. Mugele, F; Baret, J. Electrowetting: From basics to applications. J. Phys. Condens. Matter 200517, R705–R774. [Google Scholar]
  10. Quilliet, C; Berge, B. Electrowetting: A recent outbreak. Curr. Opin. Colloid Interface Sci 20016, 34–39. [Google Scholar]
  11. Probstein, RF. Physicochemical Hydrodynamics; Wiley: New York, NY, USA, 1994. [Google Scholar]
  12. Koo, J; Kleinstreuer, C. Liquid flow in microchannels: Experimental observations and computational analyses of microfluidics effects. J. Micromech. Microeng 200313, 568–579. [Google Scholar]
  13. Hu, G; Li, D. Multiscale phenomena in microfluidics and nanofluidics. Chem. Eng. Sci 200762, 3443–3454. [Google Scholar]
  14. Haus, HA; Melcher, JR. Electromagnetic Fields and Energy; Prentice-Hall: Englewood Cliffs, NJ, USA, 1989. [Google Scholar]
  15. Leal, LG. Laminar Flow and Convective Transport Processes: Scaling Principles and Asymptotic Analysis; Butterworth-Heinemann: Oxford, UK, 1992. [Google Scholar]
  16. Collins, RT; Harris, MT; Basaran, OA. Breakup of electrified jets. J. Fluid Mech 2007588, 75–129. [Google Scholar]
  17. Sista, R; Hua, Z; Thwar, P; Sudarsan, A; Srinivasan, V; Eckhardt, A; Pollack, M; Pamula, V. Development of a digital microfluidic platform for point of care testing. Lab Chip 20088, 2091–2104. [Google Scholar]
  18. Zeng, J. Modeling and simulation of electrified droplets and its application to computer-aided design of digital microfluidics. IEEE Trans. Comput. Aid. Des. Integr. Circ. Syst 200625(2), 224–233. [Google Scholar]
  19. Walker, SW; Bonito, A; Nochetto, RH. Mixed finite element method for electrowetting on dielectric with contact line pinning. Interface. Free Bound 201012, 85–119. [Google Scholar]
  20. Eck, C; Fontelos, M; Grün, G; Klingbeil, F; Vantzos, O. On a phase-field model for electrowetting. Interface. Free Bound 200911, 259–290. [Google Scholar]
  21. Gascoyne, PRC; Vykoukal, JV. Dielectrophoresis-based sample handling in general-purpose programmable diagnostic instruments. Proc. IEEE 200492(1), 22–42. [Google Scholar]
  22. Jones, TB; Gunji, M; Washizu, M. Dielectrophoretic liquid actuation and nanodroplet formation. J. Appl. Phys 200189(3), 1441–1448. [Google Scholar]
  23. Sretavan, D; Chang, W; Keller, C; Kliot, M. Microscale surgery on single axons. Neurosurgery 200557(4), 635–646. [Google Scholar]
  24. Pohl, HA; Crane, JS. Dielectrophoresis of cells. Biophys. J 197111, 711–727. [Google Scholar]
  25. Melcher, JR; Taylor, GI. Electrohydrodynamics: A review of the role of interfacial shear stresses. Annu. Rev. Fluid Mech 19691, 111–146. [Google Scholar]
  26. Saville, DA. Electrohydrodynamics: The taylor-melcher leaky-dielectric model. Annu. Rev. Fluid Mech 199729, 27–64. [Google Scholar]
  27. Schultz, GA; Corso, TN; Prosser, SJ; Zhang, S. A fully integrated monolithic microchip electrospray device for mass spectrometry. Anal. Chem 200072(17), 4058–4063. [Google Scholar]
  28. Killeen, K; Yin, H; Udiavar, S; Brennen, R; Juanitas, M; Poon, E; Sobek, D; van de Goor, T. Chip-MS: A polymeric microfluidic device with integrated mass-spectrometer interface. Micro Total Anal. Syst 2001, 331–332. [Google Scholar]
  29. Dukhin, SS. Electrokinetic phenomena of the second kind and their applications. Adv. Colloid Interface Sci 199135, 173–196. [Google Scholar]
  30. Wang, Y-C; Stevens, AL; Han, J. Million-fold preconcentration of proteins and peptides by nanofluidic filter. Anal. Chem 200577(14), 4293–4299. [Google Scholar]
  31. Kim, SJ; Wang, Y-C; Han, J. Nonlinear electrokinetic flow pattern near nanofluidic channel. Micro Total Anal. Syst 20061, 522–524. [Google Scholar]
  32. Comiskey, B; Albert, JD; Yoshizawa, H; Jacobson, J. An electrophoretic ink for all-printed reflective electronic displays. Nature 1998394(6690), 253–255. [Google Scholar]
  33. Beunis, F; Strubbe, F; Neyts, K; Bert, T; De Smet, H; Verschueren, A; Schlangen, L. P-39: Electric field compensation in electrophoretic ink display. In Proceedings of the Twenty-fifth International Display Research Conference—Eurodisplay 2005; Edinburgh, UK, 19–22 2005; pp. 344–345. [Google Scholar]
  34. Strubbe, F; Verschueren, ARM; Schlangen, LJM; Beunis, F; Neyts, K. Generation current of charged micelles in nonaqueous liquids: Measurements and simulations. J. Colloid Interface Sci 2006300, 396–403. [Google Scholar]
  35. Hsu, MF; Dufresne, ER; Weitz, DA. Charge stabilization in nonpolar solvents. Langmuir 200521, 4881–4887. [Google Scholar]
  36. Hayes, RA; Feenstra, BJ. Video-speed electronic paper based on electrowetting. Nature 2003425, 383–385. [Google Scholar]
  37. Chakrabarty, K; Su, F. Digital Microfluidic Biochips: Synthesis, Testing, and Reconfiguration Techniques; CRC Press: Boca Raton, FL, USA, 2006. [Google Scholar]
  38. Chakrabarty, K; Fair, RB; Zeng, J. Design tools for digital microfluidic biochips: Towards functional diversification and more than Moore. IEEE Trans.CAD Integr. Circ. Syst 201029(7), 1001–1017. [Google Scholar]
그림 2 : FLOW-3D를 사용한 흐름 및 형태 시뮬레이션. 파스칼 단위의 압력 및 mm 단위의 거리.

Microscopic Bubbles Switch Fiber-Optic Circuits

Figure 1: The Agilent Photonic Switching Platform
Figure 1: The Agilent Photonic Switching Platform

컴퓨터 시뮬레이션은 광섬유 회로에서 광 신호를 전환하는데 사용되는 혁신적인 스위치에서 미세 기포 문제를 이해하고 해결하는 데 중요한 역할을 했습니다. Agilent Photonic Switching Platform은 평면 광파 회로에서 잘린 작은 트렌치의 올바른 지점에 거품을 불어서 작동합니다. 버블은 광섬유 네트워크를 재구성하기 위해 광선을 다른 경로로 리디렉션 합니다. 초기 프로토타입은 기포 반사로 인해 무언가 불안정하다는 것을 나타내는 성능 문제를 보여주었습니다. 그러나 거품의 크기가 작기 때문에 문제를 진단하고 해결하는데 필요한 포괄적인 물리적 측정을 수행할 수 없었습니다.

애질런트의 선임 과학자인 John Uebbing은 전산 유체 역학 (CFD) 소프트웨어를 사용하여 거품을 시뮬레이션했습니다. 기포는 실리콘 기판에 위치한 전기 히터에 의해 유도된 증발에 의해 유지됩니다. 애질런트 팀은 트렌치 벽의 응결로 인해 유체가 축적된다는 사실을 발견했습니다. 스위치 동작의 대부분을 결정하는 것은 이러한 축적입니다. 추가 시뮬레이션을 통해 연구원들은 안정적인 신호를 제공하기 위해 장치를 변경하는 두 가지 다른 방법을 검증 할 수 있었습니다.  “처음에 우리 팀원 중 일부는 이러한 결과를 믿지 않았지만 계속된 물리적 테스트를 통해 사실이 입증되었습니다.”라고 Uebbing은 말했습니다. “CFD가 없었다면 이 문제의 해결책에 도달하지 못했을 것입니다.”

신기술 개발

광섬유 케이블은 데이터 통신 처리량을 크게 증가 시켰으며, 광 신호 전환을 위한 전기 신호로 전환한 다음 다시 광 신호로 전환하지 않고도 대량의 광섬유 데이터를 전환 할 수 있기를 원했습니다. 1990 년대 중반 Agilent Laboratories (Hewlett-Packard Labs 소속)는 전광 회로 스위치의 중요성을 인식하고 이러한 기술을 개발하기 위한 연구 프로그램을 시작했습니다. 현재 Agilent Labs의 CORL (Communications and Optical Research Laboratory) 내에 엔지니어와 과학자 팀이 구성되어 컴팩트하고 확장 가능하며 광 신호에 최소한의 영향을 미치는 이 고유한 스위치 패브릭을 개발했습니다.

 시뮬레이션은 딤플의 원인을 정확히 파악하는데 도움이 되었으며 여러 대안 솔루션을 식별하고 평가하는 데 도움이되었습니다. 버블 스위치 엔지니어링의 이러한 발전은 FLOW-3D  소프트웨어 에서 사용할 수 있는 고급 모델링 기능 없이는 불가능했을 것  입니다. 우리에게 중요한 것은 프로젝트 시작부터 Flow Science 팀이 입증한 지식과 무결성이었습니다. 우리가 이야기 한 다른 소프트웨어 회사에는 관련된 문제에 대한 표면적 이해만 있는 영업 담당자가 있었지만 Flow Science는 전문 지식을 갖춘 기술 직원을 고용하여 우리가 달성하고자 하는 것을 정확히 이해했습니다. 프로세스의 여러 단계에서 중요한 장애물을 극복 할 수 있는 중요한 도움을 제공했습니다.
– John Uebbing, 애질런트 선임 과학자

작동을 위해 Agilent Photonic Switching Platform은 두 개의 광섬유 네트워크의 교차점에 배치됩니다 (그림 1). 광 신호가 광섬유를 통해 들어 오면 직선 도파관을 통해 방해받지 않고 평면 광파 회로를 통과 할 수 있습니다. 그러나 신호가 다른 광섬유로 리디렉션되어야하는 경우 잉크젯 기술은 두 도파관 경로의 교차점에 거품을 삽입하여 광학 특성을 변경하고 신호를 출력 광섬유로의 경로 아래로 반사합니다. 기포는 거울이나 기계적으로 움직이는 부품을 사용하지 않고도 5 밀리 초 이내에 형성 및 제거 할 수 있습니다. 이 스위치는 교차된 광 도파관 배열과 인덱스 매칭되는 유체에 거품을 불어서 작동합니다. 기포는 소자 기판의 전기 히터에 의해 유도 된 증발에 의해 형성됩니다. 유체는 도파관의 교차점에 위치한 일련의 마이크로 트렌치를 채웁니다. 기포 벽으로부터의 내부 전반사로 인해 빛이 한 도파관에서 다른 도파관으로 전환됩니다. 문제는 광 도파관의 수용 각 또는 개구 수가 상당히 낮다는 것입니다. 기포의 수직 반사벽이 도파관의 축에 수직이 아니면 빛이 출력 도파관으로 제대로 반사되지 않고 신호 손실이 발생합니다.

프로토 타입의 딤플 충격 성능

초기 프로토 타입에서 광범위한 실험 테스트를 수행하여 히터 전력 및 주변 압력이 광학 반사 특성과 기포 모양 및 크기에 미치는 영향을 보여주었습니다. 이 테스트는 반사된 광 신호 대 히터 전력 곡선이 효과적인 광 스위칭에 필요한 엄격한 요구 사항을 충족하지 못하고 반사된 광 신호에 불안정성이 있음을 보여주었습니다.

그림 2 : FLOW-3D를 사용한 흐름 및 형태 시뮬레이션. 파스칼 단위의 압력 및 mm 단위의 거리.
그림 2 : FLOW-3D를 사용한 흐름 및 형태 시뮬레이션. 파스칼 단위의 압력 및 mm 단위의 거리.

컴퓨터 시뮬레이션에서 그림 2와 같이 버블의 각 면에 딤플이 형성되어 있음을 보여 주었을 때, 딤플이 전력 곡선의 혹의 원인이 되었고 반사된 신호가 그렇게 불안정한 이유 일 수 있다는 사실이 애질런트 연구팀에 나타났습니다. 센서로 물리적 측정을 수행하는 팀의 능력은 MEMS 장치의 규모까지 확장되지 않았습니다. 그들이 할 수 있는 최선은 특수 광학 장치를 사용하여 현미경 사진을 찍는 것입니다. 이 사진은 딤플이 파장 스케일에서 매우 얇기 때문에 딤플을 직접 보여줄 수 없습니다.

거품 시뮬레이션

처음에는 버블의 작동을 시뮬레이션하기 위한 여러 가지 대안이 고려되었습니다. 팀은 다양한 분석 모델을 사용하여 기포 형성을 조사했지만 이 모델은 현재 프로토 타입이 좋은 기포를 생성해야 한다고 예측했기 때문에 문제를 포착하기에는 너무 단순했습니다. 맞춤 소프트웨어를 작성하기 위해 대학 교수를 고용했지만 이 프로젝트를 완료하는 데 상당한 시간이 소요되었습니다. 그 동안 Uebbing은 문제의 복잡한 물리학을 처리 할 수 있는 상용 소프트웨어 패키지를 찾기 시작했습니다. “저는 여러 CFD 소프트웨어 개발자들과 이야기를 나눴지만 그들 중 누구도 광범위한 수정 없이 문제를 해결할 수 있는 버블 모델을 가지고 있지 않다고 판단했습니다.”라고 Uebbing은 말했습니다. “반면에 Flow Science는

Flow Science의 새로운 균질 기포 모델은 균일한 기포 압력과 온도를 가정합니다. 이것은 현실에 대한 좋은 근사치입니다. 주요 문제 중 하나는 액체, 증기 및 고체가 모두 결합되는 접점 라인의 모델링입니다. 동질 버블 모델은 이 시점에서 계산 셀의 힘과 플럭스의 균형을 맞춥니 다. Uebbing은 이전 버전의 소프트웨어를 사용하기 시작했지만 새 모델이 출시 되자마자 Uebbing은 문제를 해결해 보았습니다. “시뮬레이션 결과는 결국 실험을 설명하는 데 매우 중요한 dimple 을 보여주었습니다.”라고 Uebbing은 말했습니다. 흥미롭게도 시뮬레이션 결과 버블이 35kHz에서 진동하는 것으로 나타났습니다. 우리는 그것이 실제로 그 주파수에서 진동한다는 것을 보여주는 실험 데이터를 가져 왔지만 우리는 이유를 몰랐습니다.

현실과의 다소 예상치 못한 상관 관계는 팀에게 시뮬레이션 결과에 대한 확신을주었습니다. 시뮬레이션 결과는 문제 영역의 모든 지점에서 유속, 압력 및 온도를 보여줌으로써 테스트에서 측정 할 수 있었던 것 이상이었습니다. 이 결과로 우리는 무슨 일이 일어나고 있는지 파악할 수 있었습니다. dimple은 모세관 현상으로 인해 발생합니다. 응축액이 거품 벽에 쌓입니다. 트렌치 벽에 있는 액체의 얇은 막을 통해 빠져 나 가려고 합니다. 이러한 얇은 층을 통해 액체를 밀어 넣으려면 상당한 압력 차이가 필요합니다. 기포 벽 중앙의 높은 압력으로 인해 기포가 dimple을 형성합니다.”

문제 해결

딤플이 어떻게 형성되었는지 이해하면 안정적인 신호를 제공하기 위해 거품 모양을 수정하는 두 가지 방법이 제안되었습니다. 첫 번째는 트렌치의 유리 측벽 아래로 버블 히터를 확장하는 것입니다. 그런 다음 열이 마이크로 트렌치의 벽 위로 흘러 표면을 건조시킵니다. FLOW-3D를 사용한 시뮬레이션   은 건식 벽 거품이 매우 안정적인 스위치 신호를 제공함을 보여줍니다. 기본 물리학에 따르면 기포 온도가 벽 온도보다 낮 으면 벽이 건조해질 것입니다. 이러한 기대는 FLOW-3D  시뮬레이션 으로 확인되었습니다  .

FLOW-3D로 확인 된 두 번째 방법은 마이크로 트렌치에 소위 정적 버블을 만드는 것입니다. 장치 온도가 압력 설정 저장소 온도보다 약간 더 높으면 정적 거품이 존재합니다. 이 장치 온도는 기포를 트렌치의 모서리로 밀어 넣을 수있는 충분한 압력을 생성하지만 기포가 도파관 어레이와 히터 기판 사이의 틈을 통해 불어 나기에는 충분하지 않습니다. 이러한 정적 기포는 근처의 “crusher”기포를 사용하여 끌 수 있습니다. 이 기포는 일시적으로 충분한 과압을 생성하여 정적 기포가 붕괴되도록합니다. 분쇄기 거품 자체는 더 작은 트렌치에 있으므로 표면 장력이 작업을 완료 한 후 붕괴 될 수 있습니다. FLOW-3D 시뮬레이션은 이 모드에서 스위치 작동을 보여주기 위해 사용되었습니다.

FLOW-3D를 사용 하여 미세 유체 애플리케이션 모델링  의 성능과 다양성에 대해 자세히 알아보십시오. 

Cavitation | 캐비테이션

캐비테이션이란 무엇입니까?

The spillways of the Glen Canyon dam in 1983 (Lee and Hoopes, 1996).

캐비테이션은 유체 흐름의 매우 낮은 압력 또는 포화 압력을 높이는 온도 상승으로 인해 유체 내에서 증기 또는 기포가 빠르게 발생하는 것입니다. 기포의 갑작스런 출현 (및 후속 붕괴)은 비압축성 유체 내에서 압력의 급격한 변화를 일으켜 심각한 기계적 손상을 일으킬 수 있습니다. 캐비테이션에 의해 유도 된 힘은 1983 년 Glen Canyon 댐의 배수로에서 경험 한 손상에서 볼 수 있듯이 며칠 내에 수 피트의 암석을 침식 할 가능성이 있습니다 (Lee and Hoopes, 1996).

또한 고압 다이 캐스팅에서 캐비테이션이 발생할 수 있습니다. 다이의 수축 및 곡선을 통한 용융 합금의 빠른 이동은 급속한 압력 강하를 초래하고 후속 캐비테이션으로 이어질 수 있습니다. 생성된 증기 기포는 최종 주조에서 다공성을 유발하거나 더 나쁜 경우 다이에 손상을 일으켜 주조품을 훼손시키고 다이 수명을 감소시킬 수 있습니다.

캐비테이션은 터빈과 파이프에 손상을 줄 수 있고, 댐의 배수로에서 콘크리트를 침식하는 등의 원인이 될 수 있습니다. 아래 이미지는 댐의 배수로 바닥 근처의 콘크리트 침식을 보여줍니다. 댐에 사용되는 콘크리트는 일반적으로 강도가 높지만 캐비테이션은 여전히 그것을 부식시킬 수 있습니다.

Eroded concrete due to cavitation on the spillway of a dam

캐비테이션은 때때로 오염 물질과 유기 분자를 분해하고, 소수성 화학 물질을 결합하고, 캐비테이션 기포의 파열로 인해 생성 된 충격파를 통해 신장 결석을 파괴하고, 혼합을위한 난류를 증가시켜 수질 정화와 같은 특정 산업 응용 분야에서 의도적으로 유도됩니다.

따라서 캐비테이션이 발생할 가능성이있는 위치와 그 강도를 이해하는 것이 중요합니다. 캐비테이션을 실험을 수행하거나 실험 결과의 현상을 시각화하는 것이 어렵고, 잠재적으로 손상 될 수 있으므로 수치해석 시뮬레이션으로 검토하는 것이 매우 필요하고, 유용합니다.

Real-World Applications | 실제 응용 분야

  • 물 및 환경 구조 내에서 손상을 주는 캐비테이션 시뮬레이션
  • 다이 손상 및 주조 다공성을 유발할 수 있는 고압 다이 캐스팅 중 캐비테이션 시뮬레이션
  • MEMS 장치 내의 열 거품 형성 시뮬레이션
  • 열 전달 표면의 비등 거동 예측
  • 캐비테이션 역학으로 인한 혼합 예측

Modeling Cavitation in FLOW-3D

FLOW-3D의 캐비테이션 모델은 thermal bubble jets 와 MEMS devices를 시뮬레이션하는데 성공적으로 사용되었습니다. FLOW-3D는 “active”또는 “passive” 모델 옵션을 제공합니다. Active 모델은 기포 영역을 열고 수동 모델은 흐름을 통해 캐비테이션 기포의 존재를 추적하고 전파하지만, 기포 영역의 형성을 시작하지는 않습니다.

Active모델은 더 큰 캐비테이션 영역이 예상되고 유동장에 영향을 미치는 경우에 가장 적합하며, Passive모델은 작은 기포의 간단한 모양이 예상되는 시뮬레이션에 가장 적합합니다. 활성 모델과 에너지 전송 계산을 통해 위상 변화도 옵션입니다. 기포는 계면에서의 증발 또는 응축으로 인해 추가로 팽창하거나 수축 할 수 있습니다.

Sample Results

아래 시뮬레이션은 수축 노즐을 보여줍니다. 애니메이션은 매우 일시적인 진동 동작을 보여주는 캐비테이션 버블의 진화를 보여줍니다. 캐비테이션 부피 분율은 초기 연속 액체에서 캐비테이션의 시작을 시각화하기 위해 플롯됩니다.

아래 애니메이션은 진입 속도가 8m/s이고 수렴 기울기가 18 °이고 발산 기울기가 8 ° 인 벤츄리 내의 캐비테이션을 보여줍니다. 다시 말하지만, 캐비테이션의 과도 동작은 잘 모델링되어 있으며, 모델은 22ms의 실험 결과와 비교하여 17.4ms의 캐비테이션주기 기간을 예측합니다 (Stutz and Reboud 1997).

Cavitation in a venturi

물 탱크를 통해 이동하는 고속 발사체를 시뮬레이션하여 발사체 후류에서 생성 된 저압 영역의 공동 기둥을 보여줍니다. 발사체의 초기 속도는 600m / s입니다. 아래는 탱크의 움직임과 후행하는 캐비테이션 유체의 애니메이션입니다. 발사체가 감속함에 따라 캐비테이션 기둥의 반경이 좁아집니다.

@

High-speed bullet

References

Lee, W., Hoopes, J.A., 1996, Prediction of Cavitation Damage for Spillways, Journal of Hydraulic Engineering, 122(9): 481-488.

Plesset, M.S., Prosperetti, A., 1977, Bubble Dynamics and Cavitation, Annual Revue of Fluid Mech, 9: 145-185.

Rouse, H., 1946. Elementary Mechanics of Fluids, New York: Dover Publications, Inc.

Stutz, B., Reboud, J.L., 1997, Experiments on unsteady cavitation, Experiments in Fluids, 22: 191-198.

Micro/Biofluidics with FLOW-3D (미세/생명 유체공학)

미세/생명유체공학에 관한 모델링

  • In-Vitro Diagnostics(IVD) : 체외 진단
  • Drug Delivery : 약물 전달
  • Point of Care Devices : 현장 진료 장비
  • Microarrays : 마이크로어레이
  • Lab-on-a-chip : 랩온어칩
  • MEMS(MicroElectroMechanical Systems) : 미세전자기계시스템

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Fluid dynamics modelling for additive manufacturing

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NiTi, Ti64 및 316L에서 수행된 모델 검증

용융 풀(Melt pool) 형태 및 키홀링(Keyholing)

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Cavitation(공동현상)

Cavitation(공동현상)

공동 현상은 유체 흐름의 압력이 매우 낮거나 온도 상승으로 인해 유체 내에서 증기 및 / 또는 가스 버블이 빠르게 진화하여 포화 압력을 높입니다. 기포의 갑작스런 출현 (및 후속 붕괴)은 비압축성 유체 내에서 압력의 급격한 변화를 일으켜 심각한 기계적 손상을 일으킬 수 있습니다. 캐비테이션에 의해 유발된 힘은 1983 년 글렌 캐년 댐의 유출로에서 발생한 손상으로 볼 수 있듯이 며칠 만에 수 피트의 암석을 침식 할 가능성이 있습니다 (Lee and Hoopes, 1996).

또한, 캐비테이션은 고압 다이 캐스팅에서 발생할 수 있습니다. 다이의 수축 및 곡선을 통한 용융 합금의 빠른 이동은 압력 강하를 빠르게하여 후속 캐비테이션을 유발할 수 있습니다. 생성 된 증기 기포는 최종 주조에서 다공성을 야기하거나, 더 악화되어, 다이를 손상시켜 주조를 오염시키고 다이 수명을 감소시킬 수 있습니다. 이러한 이유로, 캐비테이션이 발생할 가능성이 있는 영역을 이해하는 것이 중요합니다. 물리적 실험을 통해 캐비테이션을 시작하고 시각화하는 것은 어렵고, 잠재적으로 피해를 주기 때문에 공정을 시뮬레이션하는 것이 바람직합니다.

실증 사례

  • 물 및 환경 구조 내에서 손상된 캐비테이션 시뮬레이션
  • 고압 다이캐스팅 중 캐비테이션을 시뮬레이션하여 다이 손상 및 캐스팅 다공성을 유발할 수 있습니다.
  • MEMS 장치 내에서 열 기포 형성 시뮬레이션
  • 열전달 표면의 비등 거동 예측
  • 캐비테이션 역학으로 인한 혼합 예측

Cavitation(공동현상) FLOW-3D모델링

FLOW-3D의 현재 캐비테이션 모델은 열 버블 제트(Thermal bubble jets) 및 MEMS 장치를 시뮬레이션하는 데 성공적으로 사용되었습니다. FLOW-3D는 “Active”또는 “Passive”모델 옵션을 제공합니다. 능동형(Active) 모델은 기포 영역을 열고 수동형 모델은 흐름을 통해 캐비테이션 기포의 존재를 추적하고 전진하지만 기포 영역의 형성은 시작하지 않습니다.

능동형(Active) 모델은 더 큰 캐비테이션 영역이 예상되고 유동장에 영향을 미치는 경우에 가장 적합한 반면, 수동형(Passive) 모델은 작은 기포의 짧은 모양이 예상되는 시뮬레이션에 가장 적합합니다. 에너지 전송의 능동 모델과 계산을 통해 위상 변화(Phase change)도 옵션입니다. 기포는 계면(Surface)에서의 증발 또는 응축으로 인해 추가로 팽창 또는 수축 될 수 있습니다.

해석 사례

아래의 결과는 8m/s의 진입 속도, 18°의 수렴 기울기 및 8°의 발산 기울기를 가진 벤투리(Venturi) 내의 캐비테이션을 보여줍니다. 캐비테이션의 과도 거동이 잘 모델링되었으며, 모델은 22ms의 실험 결과와 비교하여 17.4ms의 캐비테이션 사이클 주기(Cycle period)를 예측합니다 (Stutz and Reboud 1997).

물 탱크를 통과하는 고속 발사체를 시뮬레이트하여 발사체의 웨이크에서 발생하는 저압 영역에서 공동화 깃털(Cavitation jets)을 보여줍니다. 발사체의 초기 속도는 600m/s입니다. 아래는 탱크에서의 움직임과 후미 캐비테이션 유체의 해석 결과입니다. 캐비테이팅 플룸(Cavitating plume)의 반경은 발사체가 감속함에 따라 좁아집니다.

참고 문헌

  • Lee, W., Hoopes, J.A., 1996, Prediction of Cavitation Damage for Spillways, Journal of Hydraulic Engineering, 122(9): 481-488.
  • Plesset, M.S., Prosperetti, A., 1977, Bubble Dynamics and Cavitation, Annual Revue of Fluid Mech, 9: 145-185.
  • Rouse, H., 1946. Elementary Mechanics of Fluids, New York: Dover Publications, Inc.
  • Stutz, B., Reboud, J.L., 1997, Experiments on unsteady cavitation, Experiments in Fluids, 22: 191-198.

해석예제 및 적용사례

당사에서 오랜 기간 동안 FLOW-3D를 적용한 분야별 프로젝트 적용사례와 간단한 소개 자료를 제공합니다.
아래 분야별 적용사례 다운로드 링크를 클릭하여 자유롭게 활용하시기 바랍니다.

castinghydraulicswatermems
Casting 분야
적용사례
다운로드
Hydraulics 분야
적용사례
다운로드
WaterTreatments 분야
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다운로드
MEMS 분야
적용사례
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maritime 
Maritime 분야
적용사례
다운로드
Laser Welding 분야 적용사례 다운로드Metal 3D printing 분야 적용사례 다운로드

FLOW-3D DEM

FLOW DEM

 

FLOW DEM 은 FLOW-3D 의 기체 및 액체 유동 해석에 DEM(Discrete Element Method : 개별 요소법) 기법인 입자의 거동을 분석해주는 제품입니다.

입자 – 입자 간, 입자 – 벽 사이의 접촉이나 상호 작용을 모델링 할 수 있으므로 보다 현실적인 입자 거동의 해석이 가능합니다. 
또한 유체 부분은 전문적인 FLOW-3D 분석 기능을 사용하기 때문에 유체 와 입자거동의 연성해석을 정밀하게 또한 효율적으로 분석할 수 있습니다.

주요 기능 :
  • 고체 요소의 충돌, 스프링(Spring) / 대시 포트(Dash Pot) 모델 적용
  • Void, 1 fluid, 2 fluid(자유 계면 포함) 각각의 모드에 대응
  • 가변 밀도 / 가변 직경
  • 입자 크기조절로 입자 특성을 유지하면서 입자 수를 감소
  • 독립적인 DEM의 Sub Time Step 이용

Discrete Element Method : 개별 요소법

다수의 고체 요소의 충돌 운동을 분석하는 데 유용합니다. 유동 해석과 함께 사용하면 광범위한 용도에 응용을 할 수 있습니다.

dem1
dem2

입자 간의 충돌

Voigt model은 스프링(Spring) 및 대시 포트(Dash pot)의 조합에 의해 입자 충돌 시의 힘을 평가합니다. 탄성력 부분은 스프링 모델에서,
비탄성 충돌의 에너지 소산부분은 대시 포트 모델에서 시뮬레이션되고 있으며, 중량 및 항력은 작용하는 외력으로 고려 될 수 있습니다.

 
  • 스프링 : 변형에 관련된 힘
  • 대시 포트 : 충돌시의 상대 속도에 관련된 힘
    (점성 감쇠)
  • 스프링 및 대시 포트를 병렬로 연결
    ⇒ Voigt model
  • 힘은 법선 방향과 접선 방향으로 나누어진다

분석 모드

기본적으로 이용하는 운동 방정식은 FLOW-3D 에 사용되는 질량 입자의 운동 방정식과 같은 것이지만, 여기에 DEM으로 평가되는 항목이 추가되는 형태로 되어 있으며, 실제 시뮬레이션으로는 ‘void + DEM’, ‘1 Fluid + DEM’ , ‘ 1 Fluid 자유계면 + DEM ‘을 기본 유동 모드로 취급이 가능합니다.

dem4
dem5
dem6
void + DEM1-fluid + DEM1-fluid 자유계면 + DEM

입자 유형

입자 타입도 표준 기능의 질량 입자 모델처럼 입자 크기 (반경)와 밀도가 동일한 것 외, 크기는 같지만 밀도가 다른 것이나 밀도는 같지만 크기가 다른 것 등도 취급 가능합니다. 이로 인해 표준 질량 입자 모델에서는 입자 간의 상호 작용이 고려되어 있지 않기 때문에 모든 아래에 가라 앉아 버리고 있었지만, FLOW DEM을 이용하여 기하학적 관계를 평가하는 것이 가능합니다.

dem7-균일
-밀도 변화
-입자크기 변화

응용 분야

1. Mechanical Engineering 분야

Resin filling, screw conveyance, powder conveyance

dem8
dem9
dem10

2. Civil Engineering분야

Debris flow, gravel, falling rock

dem11
dem2

3. Chemical Engineering, Pharmaceutics 분야

Fluidized bed, cyclone, stirrer

dem12
dem13
dem14

4. MEMS, Electrical Engineering 분야

전기 입자를 포함한 전기장 해석 등

dem15

dem16

 

 

 

 

 

 

 

Coarse Graining

DEM은 일반적으로 다수의 입자를 필요로 하는 해석에 사용이 되고 있습니다. 다만 이 경우, 계산 부하가 높아지므로 현실적인 계산자원을 고려하면, 입자 수가 줄여 해석할 필요가 있습니다 .

Particle Size Increase 경우

 

중자 모래 분사 분석

DEM에서의 계산부하를 생각할 때는 입자모델에 의한 안정제한을 고려해야 하지만 서브타임스텝이라는 개념을 도입함으로써 입자의 경우와 유체의 경우의 타임스텝을 바꾸고 필요이상으로 계산시간을 들이지 않고 효율적으로 계산하는 것을 가능하게 하고 있습니다.

이를 통해 예를 들어 중자사 분사 시뮬레이션 실험에서는 이러한 문제로 자주 이용되는 빙엄 유체에서는 실험과의 정합성이 별로 좋지 않기 때문에 당사에서는 이전부터 입상류 모델이라는 모델을 개발하고 연속체로부터의 접근에서도 실험과의 높은 정합성을 실현할 수 있는 모델화를 해왔는데, 이번에 DEM을 사용해도 그것과 거의 같은 결과를 얻습니다. 할 수 있음을 확인할 수 있었다.

Reference :

  • Lefebvre D., Mackenbrock A., Vidal V., Pavan V. and Haigh PM, 2004,
  • Development and use of simulation in the Design of Blown Cores and Moulds

수치해석 용역 실적

FLOW-3D Case Studies
FLOW-3D Case Studies

수행 실적

No사업명발주처
1성남정수장 3차원 유동해석한국수자원공사
2소양강댐 홍수방지벽 설치공사 실시설계용역(수치모형실험)도화종합기술공사
3용담댐 도수터널 취수탑 유입수량 유속분포(수치모형실험)한국수자원공사
4대곡댐 여수로 문비설치 기본 및 실시설계(수치해석)도화종합기술공사
5영천댐 치수능력 증대방안 실시설계(실시모형실험)도화종합기술공사
6시화조력발전소 축조공사 턴키설계를 위한 CFD 수치모형실험대우건설
7평화의댐 2단계사업 시설공사 실시설계(수치모형실험)도화종합, 삼안건설, 한국종합개발기술공사
8광동달방댐 치수능력증대사업 기본 및 실시설계영역(수치모형실험)도화종합, 삼안건설기술공사
9광양 3단계 공업용수도 실시설계용역(여수로 수치모형실험,수어댐)삼안건설기술공사
10탐진 다목적댐 치수능력 증대방안용역(수치해석)삼안건설기술공사
11댐 상수원 설계표준도 작성용역삼안건설기술공사
12보성강댐 정밀안전진단(3D모델링 수치해석)한국시설안전관리공단
13반월정수장 노후시설 개량 기본 및 실시설계용역(수치해석 부분)한국종합엔지니어링
14청송양수발전소 1,2호기 설계기술용역/여수로 3차원 수치해석용역현대엔지니어링
15소양강댐 보조여수로 설치공사 기본설계입찰 수치모형실험용역SK건설
16잠실 수중보 어도개선 기본 및 실시설계도화종합기술공사
17서귀포시 동부하수종말처리장 고도처리시설 기본 및 실시설계용역삼안건설기술공사
18서귀포시 서부하수종말처리장 고도처리시설 기본 및 실시설계용역선진엔지니어링
19오산 제2하수처리장 건설사업입찰 기본설계용역 중 3차원 수치유동해석 분야엘지건설
20당진화력 7,8호기 취수로 수치모델링한국동서발전주식회사
21녹산배수펌프장 건설공사 대안설계용역 중 펌프장 흐름해석 부문한국종합기술개발공사
22대암댐 치수능력증대사업 기본 및 실시설계(2차) 수치해석현대엔지니어링
23용인흥덕 쓰레기 이송관로 입찰설계벽산엔지니어링
24군산하수처리장 고도처리사업 턴키공사 기본설계 전산유체해석부강테크(GS건설)
25임하댐 비상여수로 건설공사 기본설계용역(수치모형실험)삼안건설기술공사
26대청댐 비상여수로 건설공사 턴키설계용역(수치해석)삼안건설기술공사
27섬진강댐 재개발 실시설계용역(수치모형실험)삼안건설기술공사
28한강하류권급수체계구축사업 제3공구 생활용수정수장 대안설계신우엔지니어링
29임하댐 취수설비 개선공사 기본 및 실시설계용역 중 전산유체유동해석유신코퍼레이션
30광명 소하 쓰레기 자동집하시설 건설공사 T/K 기본설계용역유신코퍼레이션
31공주막여과정수장 수처리구조물의 합리적 설계를 위한 전산유체해석한국수자원공사
32김포장기지구 쓰레기 자동집하시설의 수치해석한화건설
33군장국가산단(장항지구)호안도로 축조공사 갑문수치모의실험항도엔지니어링(포스코건설)
34대청댐 비상여수로 건설공사 턴키설계용역(주)삼안
35성남판교 자동크린넷시설공사 T/K 기본설계(설계용역)건화엔지니어링
36영등포정수장 재건설 및 고도정수처리 시설공사 턴키설계용역중 수리구조물 전산 유체 해석부분삼성건설
37보령7,8호기 배수로 수치해석한국전력기술
38보령1~6호기 배수로 수치해석한국전력기술
39LNG 지하저장 실증기술개발 중 유속에 의한 Ice Ring 형성조건연구한국지질자원연구원
40LNG 지하저장 실증기술개발 중 유속에 의한 Ice Ring 형성조건연구SK건설
41파주 운정지구 쓰레기 집하시설 수집관로 수치해석건화엔지니어링
42마그네슘블록 유동,응고,응력 해석대림기업(주)
43군남홍수조절지건설공사 기본 및 실시설계용역도화종합기술공사
44안동댐 비상여수로 기본설계용역 수치모형실험에스케이건설
45세탁기 Duct 부품의 Aluminum Die-Casting CAE 해석방안 개발엘지전자
46광양 2~3연주기 고속 주조시 몰드내 열유동응고해석포스코
47Cam-shaft 다이캐스팅용 금형설계 및 주조방안 해석한국생산기술연구원
48팔당수력댐 가능최대홍수량(PMF:Probable Maximum Flood)에 의한 댐체 월류시 수리 및 구조적 안정성 검토용역한국시설안전기술공단
49담체거동을 고려한 호기조 유동해석한수테크니컬서비스
50피스톤 쿨링젯 해석기술 개발 기술용역현대자동차
51아산 방조제 배수갑문확장사업 1단계 대안설계삼안건설기술공사
52하동화력 7,8호기 냉각수 배수구 전면 저류지 축조공사 3차원 수치모형실험 해석제이슨기술단
53의암수력댐 가능최대홍수량(PMF:Probable Maximum Flood)에 의한 댐체 월류시 수리 및 구조적 안정성 검토용역한국시설안전기술공단
54춘천 및 보성강댐 가능최대홍수량(PMF:Probable Maximum Flood)에 의한 댐체 월류시 수리 및 구조적 안정성 검토용역한국시설안전기술공단
55소양강댐 여수로 방류흐름개선을 위한 수치모형실험 용역한국시설안전기술공단
56제천시 하수관거정비 임대형 민자사업(BTL) 기본설계용역 중 수충격검토(주)바셈
57금강살리기 행복지구 생태하천 조성공사계룡건설산업
58첫마을지구 생활폐기물 자동집하시설 건설공사 기본설계 T/K도화종합기술공사
59괴산댐 가능최대홍수량에 대한 댐체월류시 구조적 안정성 검토용역한국시설안전기술공단
60충남도청 이전신도시 자동집하시설 건설공사 T/K입찰 기본설계 용역(주)건화
61영등포정수장 3D 모델링(주)대우건설
62화순홍수조절지 기본 및 실시설계 용역(주)도화종합기술공사
63재천시 하수관거정비 임대형 민자사업(BTL) 기본설계용역 중 수충격검토(주)바셈
64한탄강댐본댐 및 부대시설 공사 설계 변경 용역(주)삼안
65새만금 방수제 만경5공구 건설공사 기본설계 용역(3차원 수치해석)(주)삼안
66연속 주조시 발생되는 몰드 내 열응력 영향 해석(주)엔지비
67낙동강하구둑 배수문 증설공사 기본설계용역 중3차원 수치해석(주)유신
68뚝도정수센터 시설현대화 및 고도정수처리시설 실시설계 수치해석 용역신우엔지니어링
69파주운정쓰레기 자동집하시설 건설공사(T/K)태영건설
70거제평프장도화
71광교댐수치해석도화
72Slag Pouring 및 이송 시 열유동해석매탈젠텍(POSCO)
73LICC DP매탈젠텍(POSCO)
74PFC DP 공정 해석매탈젠텍(RIST)
75행복도시하수처리장이산
76다이캐스팅 주조방안 및 해석코다코(캐스트맨 매출)
77전착성능해석용 차체모델링+전착 이차흐름현대기아기술연구소
78고열전도성 다이캐스팅 경량 방열부품개발현대자동차
79엔진/변속기1 (전륜8속 TM 케이스 및 하우징 방안설계 최적화)현대자동차
80쇽업쇼버 케이스 해석 용역현대자동차
81엔진/변속기2 (세타/실린더헤드 및 후륜 다단변속기 케이스2개 제품)현대자동차
82엔진/변속기3 / 6월현대자동차
83엔진/변속기4 / 8월현대자동차
84고강도 저밀도 산합금 열물성 DB 및 주조해석현대자동차
85진공밸브 최적화현대자동차
86Bloom 해석(연주기 몰드 내 용강 유동해석)현대제철
87상수도관망 최적관리시스템 구축사업(고성군)태성종합기술
88신월빗물저류배수시설 3차원수치해석선진ENG
89실러류 해석기술 개발현대기아기술연구소
90고덕하수처리장 수치해석그레넥스
91고덕하수처리장 수치해석엔바이로솔루션
92라오스수력발전프로젝트SK건설
93슬리브내 역비산기아차
94송석지 싸이폰 여수로농어촌공사(충남도본부 예산지사)
95고풍지 싸이폰 여수로농어촌공사(충남도본부)
96광교저수지 싸이폰 여수로지자체(수원시)
97장수지 싸이폰 여수로지자체(전남공흥군)
98광폭 마그네슘 주조기 용해로 열변형 해석용역포스코
99350톤 양수냄비 다이캐스팅 개발해피콜
100Mg 빌렛 해석HMK
101관망해석 프로그램 개발국민대학교
102충주댐 하류가물막이 수치해석대림산업
103충주댐 하류가적치 수치해석대림산업
104충주댐 하류가적치 수치해석대림산업
105평화의댐 하류부지 계획고 조정에 따른 3D 수치해석 용역대림산업
106봉화댐 실시설계 3차원 수치모형 실험도화엔지니어링
107원통수조 교반해석도화엔지니어링
108DAF 실증시설 부상조 수치해석삼진정밀
109EI과제 프로그램 개발(건기연(정우식박사))오투엔비
110SEMANGKA HEPP 수치모형 실험이산
111공릉저수지 조류 및 유속분포 유동해석한국건설기술연구원
112교육 및 해석 기술 자문한국건설기술연구원
113터빈하우징 로스트폼 주조 용역한국생산기술연구원
114터빈하우징 로스트폼 주조 용역한국생산기술연구원
115교육 및 해석 기술 자문해안해양기술
116새만금 남북2축 도로 제 3공구해석E&H컨설턴트
117달천교 교각세굴 해석E&H컨설턴트
118Lean Amine Air Cooler 부식원인 분석을 위한 유동해석GS칼텍스
119Xe Pian 하류 변경안 해석SK건설
120멤브레인 CFD 프로그램 개발국민대학교
121원형관 내부 유동해석서울시립대학교
122우수저류지 세척 시스템 해석선일엔바이로
123MD 열교환 해석(2차)알이디
124모듈조합프로그램 개발오투앤비
125해양 구조물 세굴해석전남대학교
126하우징 다이캐스팅 해석제이에스테크
127막묘듈 열교환 해석한국건설기술연구원
128두량지 PK Weir 방류량 해석한국농어촌공사
129관내 유동해석GS칼텍스
130정수장 분배수로 응집지 해석그린텍환경컨설팅
131정수장 분배수로 응집지 해석그린텍환경컨설팅
132주조제일테크
133해저구조물 세굴 및 선박유동 해석창원대학교(ADD)
134고출력 저압 램프용 자외선 반응기 해석한국건설기술연구원
135고출력 중압 램프용 자외선 반응기 해석한국건설기술연구원
136과제 해석한국건설기술연구원
137이동식보&팬스한국건설기술연구원
138Point source 기반의 하천 녹조 발생 현황 2차원 mapping 시스템한국건설기술연구원
139해석지원한국종합기술
140데이터교환customizing한국항공우주연구원
141엔진소재의 주조방안 최적화를 위한 주조해석 기술용역현대자동차
142배관유동GS건설
143울산 소수력 수치해석 용역유신
144한국건설기술연구원-이동형 해수담수화 시스템 개발 컨설팅한국건설기술연구원
145Water Dynamometer 해석두산중공업
146약액 침전 외 2건 해석세메스
147Ladle 내 Dam 및 노출부 형상변화에 따른 Vortex 거동 해석(재)포항산업과학연구원
148VMD 모듈 3D모델링알이디
149칠서정수장 기술진단 3차원 수치해석(주)그린텍환경컨설팅
150충주댐 유출부 감세지 3차원 수치해석대림산업
151친환경차용 e-4WD 유도모터 로터 주조기술개발현대자동차
152울산 #4복합 해양소수력 개발 타당성 용역중 3차원 수치해석유신
153사이펀 활용 중력구동 분리막 시스템 수치해석한국건설기술연구원
154삼척화력 소수력발전설비 설치공사(EPC) 기본 및 실시설계 중 CFD해석유신
155LG전자(평택) 생산기술원-레이저 용접 결함 예측 모델 개발LG전자(평택)
156LG전자 창원 H&A사업본부-FLOW-3D 기반 통세척 성능 해석기술 개발LG전자(창원)
 수리/수자원 분야
01 교량 설치에 따른 하천흐름 및 세굴영향 검토
컨설팅내용
  • 교량 설치로 인한 3차원 모형의 수리영향 검토
  • 세굴방지공 설치로 교량의 수리적 안정성 확보
필요데이터
  • 교각 3차원 형상 또는 도면
  • 하천 수심측량 자료 및 수치지형도
  • 하천 상/하류 홍수위 및 홍수량
해석방법
  • 하천의 유동해석 수행 후 최고유속에 해당하는 교각 선정
  • 선정교각 대상을 중심으로 세굴 모형 적용
결과물
  • 하천 유동흐름, 수위분석
  • 평형세굴심 도달시간
  • 최대세굴심 및 최대퇴적고 등
02 댐체 월류 시 수리/수문 구조적 안정성 검토
컨설팅내용
  • 상류 댐 붕괴 시 급격한 방류로 인하여 하류 댐에 미치는 영향을 검토하기 위해 댐체 월류 시 수리/수문 구조적 안정성검토
필요데이터
  • 공도교 및 수문 구조물 상세 도면
  • 하천 수심측량자료 및 주변 수치지형도
  • 하천 상/하류 홍수위 및 홍수량
해석방법
  • 상류 댐 붕괴시 홍수위/홍수량 정보입력
  • 구조물/수문 분리 후 취약한 수문 선정
  • 수문 구조해석 및 Total 힘 분석
결과물
  • 수문/구조물 받는 힘 분석
  • 굥도교 월류 여부 및 수위/유속 분포
  • 방류량 및 구조물 부압 등
 수처리 분야
01 정수처리시설 구조물 최적설계
컨설팅내용
  • 정수시설 구조물에 대한 유동, 유량, 압력, 온도분포 분석
  • 수처리과정에 발생하는 현상분석
필요데이터
  • 정수시설 구조물의 제원
  • 분배수로, 침전지 등 도면 및 3D CAD 자료
  • 초기 수위데이터 등
해석방법
  • 정수시설 구조물의 경계조건 설정
  • 형상에 따른 유동흐름 및 유량 등 초기조건 
결과물
  • 정수시설물에 작용하는 압력분포 확인
  • 유동 유입에 따른 유동양상, 유량, 유속데이터 분석
  • 온도변화에 따른 유동 및 침전효율 분석

02 하수처리시설 방류량 및 유동양상 분석
컨설팅내용
  • 토출수조의 수위 및 유동현상검토
  • 각 방류 Box의 방류유량분포 및 유속분석 
필요데이터
  • 구조물관련 설계도면 자료
  • 전체 모형 작성 및 지형데이터
  • 유체 유입량, 초기 수위관련 자료
해석방법
  • 시설 구조물에 따른 경계조건 설정
  • 초기 수위조건 및 유동현상 등 조건 확인
결과물
  • 토출 수조의 수위량 및 유동흐름
  • 유동 유입에 따른 유량, 유속데이터 분석
  • 구조물 단면의 유량흐름 데이터
 
 주조 분야
01 수축 결함최소화를 위한 주조해석
컨설팅내용
  • 주조 시 산화물 혼입방지 설계
  • 조립부 수축결함 최소화 
필요데이터
  • Frame형상 제원
  • 금형, 형상 도면자료 및 3D CAD자료
  • 초기 용탕 주입시간, 충진속도, 온도 등의 데이터
해석방법
  • 금형형상에 따른 주조해석 경계조건 설정
  • 초기 조건설정에 따른 파라미터분석
결과물
  • 충진시 산화물발생 위치 및 수축공 발생 위치
  • Solidification 확인, 결함부 현상분석
  • Gate, Runner 위치 최적화
         
02 금형 최적설계를 위한 주조해석
컨설팅내용
  • 충진 온도유지 및 제품 결함 최소화를 위한 최적설계
필요데이터
  • 금형관련 제원
  • 금형, 형상 도면자료 및 3D CAD자료
  • 초기 주조 공정조건 데이터
해석방법
  • 금형형상에 맞는 Runner, Gate 모델링
  • 용탕온도, 속도, 압력 등 조건에 따른 제품 최적설계
결과물
  • 충진시 압력분포 및 산화물 발생 위치분석
  • Solid Fraction, Solidification 등 현상분석
  • 결함부위 최소화를 위한 Gate, Runner 위치 최적화
 코팅 분야
01 Nozzle 분사를 이용한 Slit Coating 해석
컨설팅내용
  • 표면 Coating에 적합한 Nozzle 형상 설계
  • Coating 구동조건 및 압력분포 분석
필요데이터
  • 초기 Nozzle 형상 제원
  • 형상 도면자료 및 3D CAD자료
  • 초기 Coating 도포현상 및 구동조건 데이터
해석방법
  • Nozzle 구동에 따른 Coating 분석
  • 액상조건에 따른 Coating 도포형상 분석
결과물
  • Nozzle 형상 파라미터에 따른 Coating 현상분석
  • Coating 분포에 따른 높이 균일성 확인
  • 액상 온도에 따른 도포량분석
  
 MEMS 분야
01 연료전지 시스템의 최적설계를 위한 유동해석
컨설팅내용
  • 연료전지 내부형상에 따른 유동장변화 데이터
  • 유량분배에 적절한 최적의 형상조건 설계
필요데이터
  • 초기 형상 도면자료 및 3D CAD자료
  • 연료전지의 구동조건 및 물성조건
  • Actuator의 작동, 토출량, 유동 등의 데이터
해석방법
  • Micro-Channel에서의 유동분배 설정
  • 액체의 특성에 따른 토출조건 확인
결과물
  • Actuator의 속도에 따른 유동량 분석
  • Micro-Channel에서의 유동양상
  • 공동현상 최소화를 위한 최적의 구동조건

표면 장력 / Surface Tension

표면 장력 / Surface Tension

FLOW-3D에 추가 된 최초의 물리 모델 중 하나는 표면 장력이었습니다.

이 모델은 잉크젯, 무중력 환경에서의 액체 연료 거동 및 다양한 MEMS (마이크로 전자 기계 시스템) 장치와 같이 다양한 종류의 응용 분야에서 수년 동안 널리 사용되어 왔습니다. 이 후에 모델의 개선 및 확장에 대한 많은 사용자 요청이 처리되었습니다.
표면 장력에 대해 보다 나은 성능개선을 위해 FLOW-3D 버전 11에 대한 새로운 모델이 개발되었습니다. 이 모델은 계산된 모든 표면 장력의 정확성과 임의 형상의 솔리드 표면을 잡아 당기는 접착력의 정확성을 향상시킵니다. 또한 이 새로운 모델은 다공성 물질의 모세관 압력과 비 균일한 표면 장력으로 인한 접선 표면 장력을 가지고 있습니다.

새로운 모델의 예는 무중력에 포함된 원형 벽을 적시는 단순한 문제입니다.

그림 1은 실린더와 접촉각이 0 도인 물로 채워진 0.25m 직경의 실린더 75 %의 경우를 보여줍니다. 버블은 10 초 전에 벽에서 깨끗하게 분리되어 탱크를 가로 질러 움직입니다. 비 구형은 기포 표면에서 모세관 파가 전파되기 때문입니다.

그림 1. 0.0, 2.5, 5.0 및 10.0 초에 무중력에서 접촉 각이 0 인 실린더 표면의 유체 (적색) 습윤 표면.

다른 예가 그림5에 도시되어 있습니다. 2에서 서로 다른 밀도의 2 개의 초기 구형 방울이 (플롯의 색으로 표시됨) 단단한 벽을 향해 아래로 이동합니다. 플롯의 시간은 0.0, 0.01, 0.02 및 0.03 초입니다. 방울은 직경이 0.0017m, 밀도가 다르지만 표면 장력 계수는 1.872 뉴턴 / m입니다.

그림 2. 접시쪽으로 움직이는 구형의 물방울. 새로운 표면 장력 모델로 시뮬레이션. 색상은 밀도를 나타냅니다.

표면 장력 모델에 대해 자세히 알아보십시오.

Download the Flow Science Report on Surface Tension

Download Surface Tension Validation – Simple Test Problems

Applications of FLOW-3Dr to MEMS (Multiphysics Capabilities)

One of the unique features of FLOW-3Dr is the FAVORTM advantage used to accurately represent complex geometry in a rectangular Cartesian mesh.
The tedious work of generating a mesh for complex geometry is avoided. Due to this advantage, FLOW-3Dr can be used to simulate °ow in complex mi-crochannels accurately and e±ciently. Figure 1 shows a novel design of a microchannel with obliquely ori-ented or staggered herringbone ridges on its bottom wall to generate chaotic mixing of °uids (Stroock et al., Science, 295, 647-651, 2002). Fluid °ow and mix-ing through this device can be easily simulated with complicated geometry resolved using FAVOR TM. Fig-ure 2 displays the secondary °ow in two typical cross sections along the channel axis of this device. The predicted °ow ¯elds matches the experimental obser-vations of Stroock et al. very well.

 

[다운로드]

Applications of FLOW-3Dr to MEMS (Multiphysics Capabilities)

Simulating the Residue left by Evaporating Drops

Background
The “coffee ring” effect is the name given to a well known observation where the evaporative drying of a drop of coffee leaves behind a ring of dark material at the edge of the original drop. On first thought one would expect that the coffee particles, which are uniformly distributed in the drop, would simply be deposited uniformly over the area wetted by the drop. It has only been in recent years that researchers have uncovered the mechanisms that produce the ring effect (Deegan, R.D., et al).
As currently understood, the edges of drops can become pinned because of roughness or chemical elements on the surface on which they lie. Heat transfer to the drops from the substrate or the air induces evaporation, which is usually greater near the drop edge. Surface tension forces then adjust the curvature of the remaining liquid consistent with the pinned edge, which results in a net flow of liquid toward the edge. This flow replenishes the evaporative loss but also moves solute to the edge where it is concentrated by evaporation. Eventually, this mechanism builds up a ring deposit of solute at the original edge of the drop.
The residue from dried drops has implications for many useful applications, including general coating processes, formation of pixel arrays of organic materials for video displays and for a variety of micro-electro-mechanical (MEMS) devices.
Because many factors control the distribution of dried residue it is desirable to have some means to model the fluid dynamics of the process to aid engineers in making the best choices for each specific application. Such a capability has been incorporated into FLOW-3D1 making it possible to computationally investigate the influence of such parameters as the initial solute concentration, fluid viscosity, volatility of the solvent, evaporation rate, surface tension and initial shape of the drop.
This technical note presents a brief description of the residue formation model and illustrates it with several computations of an evaporating drop subject to different physical conditions.

Addition of Dielectric Phenomena to FLOW-3D

Overview
There are situations where it would be helpful to account for the interaction of electric fields with liquid and solid materials. For example, electrostatic air cleaners rely on the ability to attract small particles in flowing air to a surface where they can be collected and removed from the air. In this case the primary attractive force arises from dielectric polarization of the particles.
Spraying liquid drops onto a surface, as in spray painting, is often improved by electrifying the drops so that they repel one another and produce a more uniform distribution. Also, electrified drops can be driven to overcome air resistance by suitable electric fields.
In many types of micro-electrical-mechanical-systems (MEMS) fluids are caused to move by the application of electric potentials. Usually this behavior is induced by electric forces acting on dielectric polarization charges generated at free fluid surfaces or at the interfaces between two fluids.

In some situations the effects of both dielectrically induced charges as well as free electric charges in a fluid must be considered. For these cases the fluid has some nonzero conductivity that must be accounted for by tracking charge densities and adding additional body forces to the fluid. The range of possibilities when conduction is present includes bound and free charges, recombination, ionization, currents without net charge densities, etc. As described next, we shall limit the present development to a useful subset of the many possibilities.

In this note we describe a set of program developments that give FLOW-3DÒ the capability to model fluid and particulate flows involving both free and induced charge densities. In the current released version of FLOW-3D® (Ver. 7.7) both particles and fluid can contain a fixed charge density, but there is no provision for dielectric materials.

Here we describe the addition of dielectric properties for particles, fluids, and solids. In addition, linear polarization forces acting on particles and fluids by electrostatic fields are added to the momentum equations for fluid and particles.

Microfluidics Bibliography

Microfluidics Bibliography

다음은 Microfluidics Bibliography의 기술 문서 모음입니다.
이 모든 논문은 FLOW-3D  결과를 특징으로  합니다. 미세 유체 공정 및 장치 를 성공적으로 시뮬레이션하기 위해 FLOW-3D 를 사용 하는 방법에 대해 자세히 알아보십시오  .

2024년 3월 20일 Update

196-23 Daicong Zhang, Chunhui Jing, Wei Guo, Yuan Xiao, Jun Luo, Lehua Qi, Microchannels formed using metal microdroplets, Micromachines, 14.10; 1922, 2023. doi.org/10.3390/mi14101922

121-23 Feng Lin Ng, Zhanhong Cen, Yi-Chin Toh, Lay Poh Tan, A 3D-printed micro-perfused culture device with embedded 3D fibrous scaffold for enhanced biomimicry, International Journal of Bioprinting, 2023. doi.org/10.36922/ijb.0226

104-23 Cristina González-Fernández, Jenifer Gómez-Pastora, Eugenio Bringas, Inmaculada Ortiz, Computer-aided design of magnetophoretic microfluidic systems for enhanced recovery of target products, 33rd European Symposium on Computer-Aided Engineering (ESCAPE), 2023.

64-23   Tihomir Tjankov, Dimitar Trifonov, Conceptual design and 3D modeling of a microfluidic device for liver cells investigation, Industry 4.0, 8.2; pp. 39-41, 2023.

34-23   Chao Kang, Ikki Ikeda, Motoki Sakaguchi, Recoil and solidification of a paraffin droplet impacted on a metal substrate: Numerical study and experimental verification, Journal of Fluids and Structures, 118; 103839, 2023. doi.org/10.1016/j.jfluidstructs.2023.103839

64-22   Babatunde Aramide, Computational modelling of electrohydrodynamic jetting (Taylor cone formation, dripping & jet evolution): Case study of electrospinning, Thesis, University College London, 2022.

42-22   Islam Hassan, P. Ravi Selvaganapathy, Microfluidic printheads for highly switchable multimaterial 3D printing of soft materials, Advanced Materials Technologies, 2101709, 2022. doi.org/10.1002/admt.202101709

138-21   Enver Guler, Mine Eti, Aydin Cihanoglu, Esra Altiok, Kadriye Ozlem Hamaloglu, Burcu Gokcal, Ali Tuncel, Nalan Kabay, Ion exchange membranes with enhanced antifouling properties to produce energy from renewable sources, Proceedings of the 6th International Symposium on Green and Smart Technologies for a Sustainable Society, Santander, Cantabria, Spain, December 9-10, 2021.

45-21   Navid Tonekaboni, Mahdi Feizbahr, Nima Tonekaboni, Guang-Jun Jiang, Hong-Xia Chen, Optimization of solar CCHP systems with collector enhanced by porous media and nanofluid, Mathematical Problems in Engineering, 2021; 9984940, 2021. doi.org/10.1155/2021/9984840

40-21   B. Hayes, G.L. Whiting, R. MacCurdy, Modeling of contactless bubble–bubble interactions in microchannels with integrated inertial pumps, Physics of Fluids, 33.4; 042002, 2021. doi.org/10.1063/5.0041924

Below is a collection of technical papers in our Microfluidics Bibliography. All of these papers feature FLOW-3D results. Learn more about how FLOW-3D can be used to successfully simulate microfluidic processes and devices.

14-21   Jian-Chiun Liou, Chih-Wei Peng, Philippe Basset, Zhen-Xi Chen, DNA printing integrated multiplexer driver microelectronic mechanical system head (IDMH) and microfluidic flow estimation, Micromachines, 12.1; 25, 2021. doi.org/10.3390/mi12010025

08-20   Li Yong-Qiang, Dong Jun-Yan and Rui Wei, Numerical simulation for capillary driven flow in capsule-type vane tank with clearances under microgravity, Microgravity Science and Technology, 2020. doi.org/10.1007/s12217-019-09773-z

89-19   Tim Dreckmann, Julien Boeuf, Imke-Sonja Ludwig, Jorg Lumkemann, and Jorg Huwyler, Low volume aseptic filling: impact of pump systems on shear stress, European Journal of Pharmeceutics and Biopharmeceutics, in press, 2019. doi:10.1016/j.ejpb.2019.12.006

88-19   V. Amiri Roodan, J. Gomez-Pastora, C. Gonzalez-Fernandez, I.H. Karampelas, E. Bringas, E.P. Furlani, and I. Ortiz, CFD analysis of the generation and manipulation of ferrofluid droplets, TechConnect Briefs, pp. 182-185, 2019. TechConnect World Innovation Conference & Expo, Boston, Massachussetts, USA, June 17-19, 2019.

55-19     Julio Aleman, Sunil K. George, Samuel Herberg, Mahesh Devarasetty, Christopher D. Porada, Aleksander Skardal, and Graça Almeida‐Porada, Deconstructed microfluidic bone marrow on‐a‐chip to study normal and malignant hemopoietic cell–niche interactions, Small, 2019. doi: 10.1002/smll.201902971

37-19     Feng Lin Ng, Miniaturized 3D fibrous scaffold on stereolithography-printed microfluidic perfusion culture, Doctoral Thesis, Nanyang Technological University, Singapore, 2019.

32-19     Jenifer Gómez-Pastora, Ioannis H. Karampelas, Eugenio Bringas, Edward P. Furlani, and Inmaculada Ortiz, Numerical analysis of bead magnetophoresis from flowing blood in a continuous-flow microchannel: Implications to the bead-fluid interactions, Nature: Scientific Reports, Vol. 9, No. 7265, 2019. doi: 10.1038/s41598-019-43827-x

01-19  Jelena Dinic and Vivek Sharma, Computational analysis of self-similar capillary-driven thinning and pinch-off dynamics during dripping using the volume-of-fluid method, Physics of Fluids, Vol. 31, 2019. doi: 10.1063/1.5061715

75-18   Tobias Ladner, Sebastian Odenwald, Kevin Kerls, Gerald Zieres, Adeline Boillon and Julien Bœuf, CFD supported investigation of shear induced by bottom-mounted magnetic stirrer in monoclonal antibody formulation, Pharmaceutical Research, Vol. 35, 2018. doi: 10.1007/s11095-018-2492-4

53-18   Venoos Amiri Roodan, Jenifer Gómez-Pastora, Aditi Verma, Eugenio Bringas, Inmaculada Ortiz and Edward P. Furlani, Computational analysis of magnetic droplet generation and manipulation in microfluidic devices, Proceedings of the 5th International Conference of Fluid Flow, Heat and Mass Transfer, Niagara Falls, Canada, June 7 – 9, 2018; Paper no. 154, 2018.  doi: 10.11159/ffhmt18.154

35-18   Jenifer Gómez-Pastora, Cristina González Fernández, Marcos Fallanza, Eugenio Bringas and Inmaculada Ortiz, Flow patterns and mass transfer performance of miscible liquid-liquid flows in various microchannels: Numerical and experimental studies, Chemical Engineering Journal, vol. 344, pp. 487-497, 2018. doi: 10.1016/j.cej.2018.03.110

16-18   P. Schneider, V. Sukhotskiy, T. Siskar, L. Christie and I.H. Karampelas, Additive Manufacturing of Microfluidic Components via Wax Extrusion, Biotech, Biomaterials and Biomedical TechConnect Briefs, vol. 3, pp. 162 – 165, 2018.

15-18   J. Gómez-Pastora, I.H. Karampelas, A.Q. Alorabi, M.D. Tarn, E. Bringas, A. Iles, V.N. Paunov, N. Pamme, E.P. Furlani, I. Ortiz, CFD analysis and experimental validation of magnetic droplet generation and deflection across multilaminar flow streams, Biotech, Biomaterials and Biomedical TechConnect Briefs, vol. 3, pp. 182-185, 2018.

14-18   J. Gómez-Pastora, C. González-Fernández, I.H. Karampelas, E. Bringas, E.P. Furlani, and I. Ortiz, Design of Magnetic Blood Cleansing Microdevices through Experimentally Validated CFD Modeling, Biotech, Biomaterials and Biomedical TechConnect Briefs, vol. 3, pp. 170-173, 2018.

10-18   A. Gupta, I.H. Karampelas, J. Kitting, Numerical modeling of the formation of dynamically configurable L2 lens in a microchannel, Biotech, Biomaterials and Biomedical TechConnect Briefs, Vol. 3, pp. 186 – 189, 2018.

17-17   I.H. Karampelas, J. Gómez-Pastora, M.J. Cowan, E. Bringas, I. Ortiz and E.P. Furlani, Numerical Analysis of Acoustophoretic Discrete Particle Focusing in Microchannels, Biotech, Biomaterials and Biomedical TechConnect Briefs 2017, Vol. 3

16-17   J. Gómez-Pastora, I.H. Karampelas, E. Bringas, E.P. Furlani and I. Ortiz, CFD analysis of particle magnetophoresis in multiphase continuous-flow bioseparators, Biotech, Biomaterials and Biomedical TechConnect Briefs 2017, Vol. 3

15-17   I.H. Karampelas, S. Vader, Z. Vader, V. Sukhotskiy, A. Verma, G. Garg, M. Tong and E.P. Furlani, Drop-on-Demand 3D Metal Printing, Informatics, Electronics and Microsystems TechConnect Briefs 2017, Vol. 4

102-16   J. Brindha, RA.G. Privita Edwina, P.K. Rajesh and P.Rani, “Influence of rheological properties of protein bio-inks on printability: A simulation and validation study,” Materials Today: Proceedings, vol. 3, no.10, pp. 3285-3295, 2016. doi: 10.1016/j.matpr.2016.10.010

99-16   Ioannis H. Karampelas, Kai Liu, Fatema Alali, and Edward P. Furlani, Plasmonic Nanoframes for Photothermal Energy Conversion, J. Phys. Chem. C, 2016, 120 (13), pp 7256–7264

98-16   Jelena Dinic and Vivek Sharma, Drop formation, pinch-off dynamics and liquid transfer of simple and complex fluidshttp://meetings.aps.org/link/BAPS.2016.MAR.B53.12, APS March Meeting 2016, Volume 61, Number 2, March 14–18, 2016, Baltimore, Maryland

67-16  Vahid Bazargan and Boris Stoeber, Effect of substrate conductivity on the evaporation of small sessile droplets, PHYSICAL REVIEW E 94, 033103 (2016), doi: 10.1103/PhysRevE.94.033103

57-16   Ioannis Karampelas, Computational analysis of pulsed-laser plasmon-enhanced photothermal energy conversion and nanobubble generation in the nanoscale, PhD Dissertation: Department of Chemical and Biological Engineering, University at Buffalo, State University of New York, July 2016

44-16   Takeshi Sawada et al., Prognostic impact of circulating tumor cell detected using a novel fluidic cell microarray chip system in patients with breast cancer, EBioMedicine, Available online 27 July 2016, doi: 10.1016/j.ebiom.2016.07.027.

39-16   Chien-Hsun Wang, Ho-Lin Tsai, Yu-Che Wu and Weng-Sing Hwang, Investigation of molten metal droplet deposition and solidification for 3D printing techniques, IOP Publishing, J. Micromech. Microeng. 26 (2016) 095012 (14pp), doi: 10.1088/0960-1317/26/9/095012, July 8, 2016

30-16   Ioannis H. Karampelas, Kai Liu and Edward P. Furlani, Plasmonic Nanocages as Photothermal Transducers for Nanobubble Cancer Therapy, Nanotech 2016 Conference & Expo, May 22-25, Washington, DC.

29-16   Scott Vader, Zachary Vader, Ioannis H. Karampelas and Edward P. Furlani, Advances in Magnetohydrodynamic Liquid Metal Jet Printing, Nanotech 2016 Conference & Expo, May 22-25, Washington, DC.

02-16  Stephen D. Hoath (Editor), Fundamentals of Inkjet Printing: The Science of Inkjet and Droplets, ISBN: 978-3-527-33785-9, 472 pages, February 2016 (see chapters 2 and 3 for FLOW-3D results)

125-15   J. Berthier, K.A. Brakke, E.P. Furlani, I.H. Karampelas, V. Poher, D. Gosselin, M. Cubinzolles and P. Pouteau, Whole blood spontaneous capillary flow in narrow V-groove microchannels, Sensors and Actuators B: Chemical, 206, pp. 258-267, 2015.

86-15   Yousub Lee and Dave F. Farson, Simulation of transport phenomena and melt pool shape for multiple layer additive manufacturing, J. Laser Appl. 28, 012006 (2016). doi: 10.2351/1.4935711, published online 2015.

77-15   Ho-Lin Tsai, Weng-Sing Hwang, Jhih-Kai Wang, Wen-Chih Peng and Shin-Hau Chen, Fabrication of Microdots Using Piezoelectric Dispensing Technique for Viscous Fluids, Materials 2015, 8(10), 7006-7016. doi: 10.3390/ma8105355

63-15   Scott Vader, Zachary Vader, Ioannis H. Karampelas and Edward P. Furlani, Magnetohydrodynamic Liquid Metal Jet Printing, TechConnect World Innovation Conference & Expo, Washington, D.C., June 14-17, 2015

46-15   Adwaith Gupta, 3D Printing Multi-Material, Single Printhead Simulation, Advanced Qualification of Additive Manufacturing Materials Workshop, July 20 – 21, 2015, Santa Fe, NM

28-15   Yongqiang Li, Mingzhu Hu, Ling Liu, Yin-Yin Su, Li Duan, and Qi Kang, Study of Capillary Driven Flow in an Interior Corner of Rounded Wall Under MicrogravityMicrogravity Science and Technology, June 2015

20-15   Pamela J. Waterman, Diversity in Medical Simulation Applications, Desktop Engineering, May 2015, pp 22-26,

16-15   Saurabh Singh, Ann Junghans, Erik Watkins, Yash Kapoor, Ryan Toomey, and Jaroslaw Majewski, Effects of Fluid Shear Stress on Polyelectrolyte Multilayers by Neutron Scattering Studies, © 2015 American Chemical Society, DOI: 10.1021/acs.langmuir.5b00037, Langmuir 2015, 31, 2870−2878, February 17, 2015

11-15   Cheng-Han Wu and Weng-Sing Hwang, The effect of process condition of the ink-jet printing process on the molten metallic droplet formation through the analysis of fluid propagation direction, Canadian Journal of Physics, 2015. doi: 10.1139/cjp-2014-0259

03-15 Hanchul Cho, Sivasubramanian Somu, Jin Young Lee, Hobin Jeong and Ahmed Busnaina, High-Rate Nanoscale Offset Printing Process Using Directed Assembly and Transfer of Nanomaterials, Adv. Materials, doi: 10.1002/adma.201404769, February 2015

122-14  Albert Chi, Sebastian Curi, Kevin Clayton, David Luciano, Kameron Klauber, Alfredo Alexander-Katz, Sebastián D’hers and Noel M Elman, Rapid Reconstitution Packages (RRPs) implemented by integration of computational fluid dynamics (CFD) and 3D printed microfluidics, Research Gate, doi: 10.1007/s13346-014-0198-7, July 2014

113-14 Cihan Yilmaz, Arif E. Cetin, Georgia Goutzamanidis, Jun Huang, Sivasubramanian Somu, Hatice Altug, Dongguang Wei and Ahmed Busnaina, Three-Dimensional Crystalline and Homogeneous Metallic Nanostructures Using Directed Assembly of Nanoparticles, 10.1021/nn500084g, © 2014 American Chemical Society, April 2014

110-14 Koushik Ponnuru, Jincheng Wu, Preeti Ashok, Emmanuel S. Tzanakakis and Edward P. Furlani, Analysis of Stem Cell Culture Performance in a Microcarrier Bioreactor System, Nanotech, Washington, D.C., June 15-18, 2014

109-14   Ioannis H. Karampelas, Young Hwa Kim and Edward P. Furlani, Numerical Analysis of Laser Induced Photothermal Effects using Colloidal Plasmonic Nanostructures, Nanotech, Washington, D.C., June 15-18, 2014

108-14   Chenxu Liu, Xiaozheng Xue and Edward P. Furlani, Numerical Analysis of Fully-Coupled Particle-Fluid Transport and Free-Flow Magnetophoretic Sorting in Microfluidic Systems, Nanotech, Washington, D.C., June 15-18, 2014

95-14   Cheng-Han Wu, Weng-Sing Hwang, The effect of the echo-time of a bipolar pulse waveform on molten metallic droplet formation by squeeze mode piezoelectric inkjet printing, Accepted November 2014, Microelectronics Reliability (2014) , © 2014 Elsevier Ltd. All rights reserved.

85-14   Sudhir Srivastava, Lattice Boltzmann method for contact line dynamics, ISBN: 978-90-386-3608-5, Copyright © 2014 S. Srivastava

61-14   Chenxu Liu, A Computational Model for Predicting Fully-Coupled Particle-Fluid Dynamics and Self-Assembly for Magnetic Particle Applications, Master’s Thesis: State University of New York at Buffalo, 2014, 75 pages; 1561583, http://gradworks.umi.com/15/61/1561583.html

41-14 Albert Chi, Sebastian Curi, Kevin Clayton, David Luciano, Kameron Klauber, Alfredo Alexander-Katz, Sebastian D’hers, and Noel M. Elman, Rapid Reconstitution Packages (RRPs) implemented by integration of computational fluid dynamics (CFD) and 3D printed microfluidics, Drug Deliv. and Transl. Res., DOI 10.1007/s13346-014-0198-7, # Controlled Release Society 2014. Available for purchase online at SpringerLink.

21-14  Suk-Hee Park, Ung Hyun Koh, Mina Kim, Dong-Yol Yang, Kahp-Yang Suh and Jennifer Hyunjong Shin, Hierarchical multilayer assembly of an ordered nanofibrous scaffold via thermal fusion bonding, Biofabrication 6 (2014) 024107 (10pp), doi:10.1088/1758-5082/6/2/024107, IOP Publishing, 2014. Available for purchase online at IOP.

17-14   Vahid Bazargan, Effect of substrate cooling and droplet shape and composition on the droplet evaporation and the deposition of particles, Ph.D. Thesis: Department of Mechanical Engineering, The University of British Columbia, March 2014, © Vahid Bazargan, 2014

73-13  Oliver G. Harlen, J. Rafael Castrejón-Pita, and Arturo Castrejon-Pita, Asymmetric Detachment from Angled Nozzles Plates in Drop-on Demand Inkjet Printing, NIP & Digital Fabrication Conference, 2013 International Conference on Digital Printing Technologies. Pages 253-549, pp. 277-280(4)

63-13  Fatema Alali, Ioannis H. Karampelas, Young Hwa Kim, and Edward P. Furlani, Photonic and Thermofluidic Analysis of Colloidal Plasmonic Nanorings and Nanotori for Pulsed-Laser Photothermal ApplicationsJ. Phys. Chem. C, Article ASAP, DOI: 10.1021/jp406986y, Copyright © 2013 American Chemical Society, September 2013.

25-13  Sudhir Srivastava, Theo Driessen, Roger Jeurissen, Herma Wijshoff, and Federico Toschi, Lattice Boltzmann Method to Study the Contraction of a Viscous Ligament, International Journal of Modern Physics © World Scientific Publishing Company, May 2013.

11-13  Li-Chieh Hsu, Yong-Jhih Chen, Jia-Huang Liou, Numerical Investigation in the Factors on the Pool Boiling, Applied Mechanics and Materials Vol. 311 (2013) pp 456-461, © (2013) Trans Tech Publications, Switzerland, doi:10.4028/www.scientific.net/AMM.311.456. Available for purchase online at Scientific.Net.

10-13 Pamela J. Waterman, CFD: Shaping the Medical World, Desktop Engineering, April 2013. Full article available online at Desktop Engineering.

90-12 Charles R. Ortloff and Martin Vogel, Spray Cooling Heat Transfer- Test and CFD Analysis, Electronics Cooling, June 2012. Available online at Electronics Cooling.

79-12    Daniel Parsaoran Siregar, Numerical simulation of evaporation and absorption of inkjet printed droplets, Ph.D. Thesis: Technische Universiteit Eindhoven, September 18, 2012, Copyright 2012 by D.P. Siregar, ISBN: 978-90-386-3190-5.

71-12   Jong-hyeon Chang, Kyu-Dong Jung, Eunsung Lee, Minseog Choi, Seungwan Lee, and Woonbae Kim, Varifocal liquid lens based on microelectrofluidic technology, Optics Letters, Vol. 37, Issue 21, pp. 4377-4379 (2012) http://dx.doi.org/10.1364/OL.37.004377

70-12   Jong-hyeon Chang, Kyu-Dong Jung, Eunsung Lee, Minseog Choi, and Seunwan Lee, Microelectrofluidic Iris for Variable ApertureProc. SPIE 8252, MOEMS and Miniaturized Systems XI, 82520O (February 9, 2012); doi:10.1117/12.906587

69-12   Jong-hyeon Chang, Eunsung Lee, Kyu-Dong Jung, Seungwan Lee, Minseog Choi, and  Woonbae Kim, Microelectrofluidic Lens for Variable CurvatureProc. SPIE 8486, Current Developments in Lens Design and Optical Engineering XIII, 84860X (October 11, 2012); doi:10.1117/12.925852.

61-12  Biddut Bhattacharjee, Study of Droplet Splitting in an Electrowetting Based Digital Microfluidic System, Thesis: Doctor of Philosophy in the College of Graduate Studies (Applied Sciences), The University of British Columbia, September 2012, © Biddut Bhattacharjee.

55-12 Hejun Li, Pengyun Wang, Lehua Qi, Hansong Zuo, Songyi Zhong, Xianghui Hou, 3D numerical simulation of successive deposition of uniform molten Al droplets on a moving substrate and experimental validation, Computational Materials Science, Volume 65, December 2012, Pages 291–301. Available for purchase online at SciVerse.

54-12   Edward P. Furlani, Anthony Nunez, Gianmarco Vizzeri, Modeling Fluid Structure-Interactions for Biomechanical Analysis of the Human Eye, Nanotech Conference & Expo, June 18-21, 2012, Santa Clara, CA.

53-12   Xinyun Wu, Richard D. Oleschuk and Natalie M. Cann, Characterization of microstructured fibre emitters in pursuit of improved nano electrospray ionization performance, The Royal Society of Chemistry 2012, http://pubs.rsc.org, DOI: 10.1039/c2an35249d, May 2012

25-12    Edward P. Furlani, Ioannis H. Karampelas and Qian Xie, Analysis of Pulsed Laser Plasmon-assisted Photothermal Heating and Bubble Generation at the Nanoscale, Lab on a Chip, 10.1039/C2LC40495H, Received 01 May 2012, Accepted 07 Jun 2012. First published on the web 13 Jun 2012.

22-12  R.A. Sultanov, D. Guster, Numerical Modeling and Simulations of Pulsatile Human Blood Flow in Different 3D-Geometries, Book chapter #21 in Fluid Dynamics, Computational Modeling and Applications (2012), ISBN: 978-953-51-0052-2, p. 475 [18 pages]. Available online at INTECH.

21-12  Guo-Wei Huang, Tzu-Yi Hung, and Chin-Tai Chen, Design, Simulation, and Verification of Fluidic Light-Guide Chips with Various Geometries of Micro Polymer Channels, NEMS 2012, Kyoto, Japan, March 5-8, 2012. Available for purchase online at IEEE.

103-11   Suk-Hee Park, Development of Three-Dimensional Scaffolds containing Electrospun Nanofibers and their Applications to Tissue Regeneration, Ph.D. Thesis: School of Mechanical, Aersospace and Systems Engineering, Division of Mechanical Engineering, KAIST, 2011.

81-11   Xinyun Wu, Modeling and Characterization of Microfabricated Emitters-In Pursuit of Improved ESI-MS Performance, thesis: Department of Chemistry, Queen’s University, December 2011, Copyright © Xinyun Wu, 2011

79-11  Cong Lu, A Cell Preparation Stage for Automatic Cell Injection, thesis: Graduate Department of Mechanical and Industrial Engineering, University of Toronto, Copyright © Cong Lu, 2011

77-11 Ge Bai, W. Thomas Leach, Computational fluid dynamics (CFD) insights into agitation stress methods in biopharmaceutical development, International Journal of Pharmaceutics, Available online 8 December 2011, ISSN 0378-5173, 10.1016/j.ijpharm.2011.11.044. Available online at SciVerse.

72-11  M.R. Barkhudarov, C.W. Hirt, D. Milano, and G. Wei, Comments on a Comparison of CFD Software for Microfluidic Applications, Flow Science Technical Note #93, FSI-11-TN93, December 2011

45-11  Chang-Wei Kang, Jiak Kwang Tan, Lunsheng Pan, Cheng Yee Low and Ahmed Jaffar, Numerical and experimental investigations of splat geometric characteristics during oblique impact of plasma spraying, Applied Surface Science, In Press, Corrected Proof, Available online 20 July 2011, ISSN 0169-4332, DOI: 10.1016/j.apsusc.2011.06.081. Available to purchase online at SciVers

33-11  Edward P. Furlani, Mark T. Swihart, Natalia Litchinitser, Christopher N. Delametter and Melissa Carter, Modeling Nanoscale Plasmon-assisted Bubble Nucleation and Applications, Nanotech Conference and Expo 2011, Boston, MA, June 13-16, 2011

32-11  Lu, Cong and Mills, James K., Three cell separation design for realizing automatic cell injection, Complex Medical Engineering (CME), 2011 IEEE/ICME, pp: 599 – 603, Harbin, China, 10.1109/ICCME.2011.5876811, June 2011. Available online at IEEEXplore.

25-11 Issam M. Bahadur, James K. Mills, Fluidic vacuum-based biological cell holding device with piezoelectrically induced vibration, Complex Medical Engineering (CME), 2011 IEEE/ICME International Conference on, 22-25 May 2011, pp: 85 – 90, Harbin, China. Available online at: IEEE Xplore.

14-11  Edward P. Furlani, Roshni Biswas, Alexander N. Cartwright and Natalia M. Litchinitser, Antiresonant guiding optofluidic biosensor, doi:10.1016/j.optcom.2011.04.014, Optics Communication, April 2011

05-11 Hyeju Eom and Keun Park, Integrated numerical analysis to evaluate replication characteristics of micro channels in a locally heated mold by selective induction, International Journal of Precision Engineering and Manufacturing, Volume 12, Number 1, 53-60, DOI: 10.1007/s12541-011-0007-x, 2011. Available online at: SpringerLink.

70-10  I.N. Volnov, V.S. Nagornyi, Modeling Processes for Generation of Streams of Monodispersed Fluid Droplets in Electro-inkjet Applications, Science and Technology News, St. Petersburg State Polytechnic University, 4, pp 294-300, 2010. In Russian.

62-10  F. Mobadersani, M. Eskandarzade, S. Azizi and S. Abbasnezhad, Effect of Ambient Pressure on Bubble Growth in Micro-Channel and Its Pumping Effect, ESDA2010-24436, pp. 577-584, doi:10.1115/ESDA2010-24436, ASME 2010 10th Biennial Conference on Engineering Systems Design and Analysis (ESDA2010), Istanbul, Turkey, July 12–14, 2010. Available online at the ASME Digital Library.

58-10 Tsung-Yi Ho, Jun Zeng, and Chakrabarty, K, Digital microfluidic biochips: A vision for functional diversity and more than moore, Computer-Aided Design (ICCAD), 2010 IEEE/ACM International Conference on, DOI: 10.1109/ICCAD.2010.5654199, © IEEE, November 2010. Available online at IEEE Explore.

51-10  Regina Bleul, Marion Ritzi-Lehnert, Julian Höth, Nico Scharpfenecker, Ines Frese, Dominik Düchs, Sabine Brunklaus, Thomas E. Hansen-Hagge, Franz-Josef Meyer-Almes, Klaus S. Drese, Compact, cost-efficient microfluidics-based stopped-flow device, Anal Bioanal Chem, DOI 10.1007/s00216-010-4446-5, Available online at Springer, November 2010

22-10    Krishendu Chakrabarty, Richard B. Fair and Jun Zeng, Design Tools for Digital Microfluidic Biochips Toward Functional Diversification and More than Moore, IEEE Transactions on Computer-Aided Design of Integrated Circuits and Systems, Vol. 29, No. 7, July 2010

14-10 E. P. Furlani and M. S. Hanchak, Nonlinear analysis of the deformation and breakup of viscous microjets using the method of lines, International Journal for Numerical Methods in Fluids (2010), © 2010 John Wiley & Sons, Ltd., Published online in Wiley InterScience. DOI: 10.1002/fld.2205

55-09 R.A. Sultanov, and D. Guster, Computer simulations of  pulsatile human blood flow through 3D models of the human aortic arch, vessels of simple geometry and a bifurcated artery, Proceedings of the 31st Annual International Conference of the IEEE EMBS (Engineering in Medicine and Biology Society), Minneapolis, September 2-6, 2009, p.p. 4704-4710.

30-09 Anurag Chandorkar and Shayan Palit, Simulation of Droplet Dynamics and Mixing in Microfluidic Devices using a VOF-Based Method, Sensors & Transducers journal, ISSN 1726-5479 © 2009 by IFSA, Vol.7, Special Issue “MEMS: From Micro Devices to Wireless Systems,” October 2009, pp. 136-149.

13-09 E.P. Furlani, M.C. Carter, Analysis of an Electrostatically Actuated MEMS Drop Ejector, Presented at Nanotech Conference & Expo 2009, Houston, Texas, USA, May 3-7, 2009

12-09 A. Chandorkar, S. Palit, Simulation of Droplet-Based Microfluidics Devices Using a Volume-of-Fluid Approach, Presented at Nanotech Conference & Expo 2009, Houston, Texas, USA, May 3-7, 2009

3-09 Christopher N. Delametter, FLOW-3D Speeds MEMS Inkjet Development, Desktop Engineering, January 2009

42-08  Tien-Li Chang, Jung-Chang Wang, Chun-Chi Chen, Ya-Wei Lee, Ta-Hsin Chou, A non-fluorine mold release agent for Ni stamp in nanoimprint process, Microelectronic Engineering 85 (2008) 1608–1612

26-08 Pamela J. Waterman, First-Pass CFD Analyses – Part 2, Desktop Engineering, November 2008

09-08 M. Ren and H. Wijshoff, Thermal effect on the penetration of an ink droplet onto a porous medium, Proc. Eurotherm2008 MNH, 1 (2008)

04-08 Delametter, Christopher N., MEMS development in less than half the time, Small Times, Online Edition, May 2008

02-08 Renat A. Sultanov, Dennis Guster, Brent Engelbrekt and Richard Blankenbecler, 3D Computer Simulations of Pulsatile Human Blood Flows in Vessels and in the Aortic Arch – Investigation of Non-Newtonian Characteristics of Human Blood, The Journal of Computational Physics, arXiv:0802.2362v1 [physics.comp-ph], February 2008

01-08 Herman Wijshoff, thesis: University of Twente, Structure- and fluid dynamics in piezo inkjet printheads, ISBN 978-90-365-2582-4, Venlo, The Netherlands January 2008.

30-07 A. K. Sen, J. Darabi, and D. R. Knapp, Simulation and parametric study of a novel multi-spray emitter for ESI–MS applications, Microfluidics and Nanofluidics, Volume 3, Number 3, June 2007, pp. 283-298(16)

28-07 Dan Soltman and Vivek Subramanian, Inkjet-Printed Line Morphologies and Temperature Control of the Coffee Ring Effect, Langmuir; 2008; ASAP Web Release Date: 16-Jan-2008; (Research Article) DOI: 10.1021/la7026847

23-07 A K Sen and J Darabi, Droplet ejection performance of a monolithic thermal inkjet print head, Journal of Micromechanical and Microengineering,vol.17, pp.1420-1427 (2007) doi:10.1088/0960-1317/17/8/002; Abstract only.

18-07 Herman Wisjhoff, Better Printheads Via Simulation, Desktop Engineering, October 2007, Vol. 13, Issue 2

17-07 Jos de Jong, Ph.D. Thesis: University of Twente, Air entrapment in piezo inkjet printing, ISBN 978-90-365-2483-4, April 2007

15-07 Krishnendu Chakrabarty and Jun Zeng, (Ed.), Design Automation Methods and Tools for Microfluidics-Based Biochips, Springer, September 2006.

14-07 Fei Su and Jun Zeng, Computer-aided design and test for digital microfluidics, IEEE Design & Test of Computers, 24(1), 2007, 60-70.

13-07 Jun Zeng, Modeling and simulation of electrified droplets and its application to computer-aided design of digital microfluidics, IEEE Transactions on Computer-Aided Design of Integrated Circuits and Systems, 25(2), 2006, 224-233.

12-07 Krishnendu Chakrabarty and Jun Zeng, (2005), Automated top-down design for microfluidic biochips, ACM Journal on Emerging Technologies in Computing Systems, 1(3), 2005, 186–223.

01-07 Wijshoff, Herman, Drop formation mechanisms in piezo-acoustic inkjet, NSTI-Nanotech 2007, ISBN 1420061844 Vol. 3, 2007)

23-06 John J. Uebbing, Stephan Hengstler, Dale Schroeder, Shalini Venkatesh, and Rick Haven, Heat and Fluid Flow in an Optical Switch Bubble, Journal of Microelectromechanical Systems, Vol. 15, No. 6, December 2006

21-06 Wijshoff, Herman, Manipulating Drop Formation in Piezo Acoustic Inkjet, Proc. IS&T’s NIP22, 79 (2006)

20-06 J. de Jong, H. Reinten, M. van den Berg, H. Wijshoff, M. Versluis, G. de Bruin, A. Prosperetti and D. Lohse, Air entrapment in piezo-driven inkjet printheads, J. Acoust. Soc. Am. 120(3), 1257 (2006)

11-06 A. K. Sen, J. Darabi, D. R. Knapp and J. Liu, Modeling and Characterization of a Carbon Fiber Emitter for Electrospray Ionization, 1 MEMS and Microsystems Laboratory, Department of Mechanical Engineering, University of South Carolina, 300 Main Street, Columbia, SC 29208, USA, 2 Department of Pharmacology, Medical University of South Carolina, Charleston, SC

5-06 E. P. Furlani, B. G. Price, G. Hawkins, and A. G. Lopez, Thermally Induced Marangoni Instability of Liquid Microjets with Application to Continuous Inkjet Printing, Proceedings of NSTI Nanotech Conference 2006, Vol. 2, pp 534-537.

28-05 O B Fawehinmi, P H Gaskell, P K Jimack, N Kapur, and H M Thompson, A combined experimental and computational fluid dynamics analysis of the dynamics of drop formation, May 2005. DOI: 10.1243/095440605X31788

5-05 E. P. Furlani, Thermal Modulation and Instability of Newtonian Liquid Microjets, presented at Nanotech 2005, Anaheim, CA, May 8-12, 2005.

1-05 C.W. Hirt, Electro-Hydrodynamics of Semi-Conductive Fluids: With Application to Electro-Spraying, Flow Science Technical Note #70, FSI-05-TN70

19-04 G. F. Yao, Modeling of Electroosmosis Without Resolving Physics Inside a Electric Double Layer, Flow Science Technical Note (FSI-04-TN69)

12-04 Jun Zeng and Tom Korsmeyer, Principles of Droplet Electrohydrodynamics for Lab-on-a-Chip, Lab. Chip. Journal, 2004, 4(4), 265-277

9-04 Constantine N. Anagnostopoulos, James M. Chwalek, Christopher N. Delametter, Gilbert A. Hawkins, David L. Jeanmaire, John A. Lebens, Ali Lopez, and David P. Trauernicht, Micro-Jet Nozzle Array for Precise Droplet Metering and Steering Having Increased Droplet Deflection, Proceedings of the 12th International Conference on Solid State Sensors, Actuators and Microsystems, sponsored by IEEE, Boston, June 8-12, 2003, pp. 368-71

8-04 Christopher N. Delametter, David P. Trauernicht, James M. Chwalek, Novel Microfluidic Jet Deflection – Significant Modeling Challenge with Great Application Potential, Technical Proceedings of the 2002 International Conference on Modeling and Simulation of Microsystems sponsored by NSTI, San Juan, Puerto Rico, April 21-25, 2002, pp. 44-47

6-04 D. Vadillo*, G. Desie**, A Soucemarianadin*, Spreading Behavior of Single and Multiple Drops, *Laboratoire des Ecoulements Geophysiques et Industriels (LEGI), and **AGFA-Gevaert Group N.V., XXI ICTAM, 15-21 August 2004, Warsaw, Poland

2-04 Herman Wijshoff, Free Surface Flow and Acousto-Elastic Interaction in Piezo Inkjet, Nanotech 2004, sponsored by the Nano Science & Technology Institute, Boston, MA, March 2004

30-03 D Souders, I Khan and GF Yao, Alessandro Incognito, and Matteo Corrado, A Numerical Model for Simulation of Combined Electroosmotic and Pressure Driven Flow in Microdevices, 7th International Symposium on Fluid Control, Measurement and Visualization

27-03 Jun Zeng, Daniel Sobek and Tom Korsmeyer, Electro-Hydrodynamic Modeling of Electrospray Ionization – CAD for a µFluidic Device-Mass Spectrometer Interface, Agilent Technologies Inc, paper presented at Transducers 2003, June 03 Boston (note: Reference #10 is to FLOW-3D)

17-03 John Uebbing, Switching Fiber-optic Circuits with Microscopic Bubbles, Sensors Magazine, May 2003, Vol 20, No 5, p 36-42

16-03 CFD Speeds Development of MEMS-based Printing Technology, MicroNano Magazine, June 2003, Vol 8, No 6, p 16

3-03 Simulation Speeds Design of Microfluidic Medical Devices, R&D Magazine, March 2003, pp 18-19

1-03 Simulations Help Microscopic Bubbles Switch Fiber-Optic Circuits, Agilent Technologies, Fiberoptic Product News, January 2003, pp 22-23

27-02 Feng, James Q., A General Fluid Dynamic Analysis of Drop Ejection in Drop-on-Demand Ink Jet Devices, Journal of Imaging Science and Technology®, Volume 46, Number 5, September/October 2002

1-02 Feixia Pan, Joel Kubby, and Jingkuang Chen, Numerical Simulation of Fluid Structure Interaction in a MEMS Diaphragm Drop Ejector, Xerox Wilson Research Center, Institute of Physics Publishing, Journal of Micromechanics and Microengineering, 12 (2002), PII: SO960-1317(02)27439-2, pp. 70-76

48-01   Rainer Gruber, Radial Mass Transfer Enhancement in Bubble-Train Flow, PhD thesis in Engineering Sciences, Rheinisch- Westf alischen Technische Hochschule Aachen, December 2001.

34-01 Furlani, E.P., Delametter, C.N., Chwalek, J.M., and Trauernicht, D., Surface Tension Induced Instability of Viscous Liquid Jets, Fourth International Conference on Modeling and Simulation of Microsystems, April 2001

12-01 C. N. Delametter, Eastman Kodak Company, Micro Resolution, Mechanical Engineering, Col 123/No 7, July 2001, pp 70-72

11-01 C. N. Delametter, Eastman Kodak Company, Surface Tension Induced Instability of Viscous Liquid Jets, Technical Proceeding of the Fourth International Conference on Modeling and Simulation of Microsystems, April 2001

9-01 Aman Khan, Unipath Limited Research and Development, Effects of Reynolds Number on Surface Rolling in Small Drops, PVP-Col 431, Emerging Technologies for Fluids, Structures and Fluids, Structures and Fluid Structure Interaction — 2001

2-00 Narayan V. Deshpande, Significance of Inertance and Resistance in Fluidics of Thermal Ink-Jet Transducers, Journal of Imaging Science and Technology, Volume 40, Number 5, Sept./Oct. 1996, pp.457-461

4-98 D. Deitz, Connecting the Dots with CFD, Mechanical Engineering Magazine, pp. 90-91, March 1998

14-94 M. P. O’Hare, N. V. Deshpande, and D. J. Drake, Drop Generation Processes in TIJ Printheads, Xerox Corporation, Adv. Imaging Business Unit, IS&T’s Tenth International Congress on Advances in Non-Impact Printing, Tech. 1994

14-92 Asai, A.,Three-Dimensional Calculation of Bubble Growth and Drop Ejection in a Bubble Jet Printer, Journal of Fluids Engineering Vol. 114 December 1992:638-641

MEMS/WELD 분야

Microfluidics

Microfluidics는 집적 회로 산업에서 사용되는 것과 유사한 공정을 사용하여 소형 기기의 제조에 급격하게 성장하는 기술입니다. Microfluidics 기술은 0.1 미크론에서 1mm에 이르기까지 매우 작은 장치로 기계, 유체, 광학, 전자 기능을 통합 할 수있는 방법을 제공합니다. Microfluidics는 기존의 방법과 비교하면 두 가지 중요한 장점이 있습니다. 첫째, 대량으로 제조 될 수 있으므로, 생산의 비용이 실질적으로 감소 될 수 있습니다. 둘째, 집적 회로에 통합 될 수 있어서 다른 기술보다 훨씬 더 복잡한 시스템으로 제조 될 수 있습니다.

Chip packaging simulation. Results generated by FLOW-3D/MP, FLOW-3D‘s HPC solution.

엔지니어 및 과학자가 설계, 시험 제작하고 그 성능을 최적화하기 위해 장치를 재 설계하는 등, 다른 제조 방법에서와 같이 microfluidics 설계 프로세스는 매우 고가 일 수 있습니다. 그러나, 수치 시뮬레이션은 전자, 기계, 화학, 열 과학 및 유체 과학 등의 분야에 걸쳐 정량 분석과 중요한 통찰력을 제공 할 수 있습니다.

laser-sintering

 

분야별 적용사례

  • 수자원 공학 분야 : spillways, 댐 붕괴, 홍수 영향, 쓰나미, 수처리 및 하수 시스템을 통한 물 흐름을 추적하고 분석하는 분야에서 다양하게 활용되고 있습니다.
  • 항공 우주 분야 : FLOW-3D 사용하여 항공기의 연료 탱크, 특히 우주선의 연료 유출을 정확하게 예측합니다. 
  • 적층 제조 분야 :  FLOW-3D 는 적층 제조분야에서 현업에서 사용할 수 있는 유일한 CFD소프트웨어 중 하나로, 파우더 베드 퓨전과 같은 다양한 정교한 적층 제조 프로세스를 상세한 수준으로 모델링 할 수 있습니다.
  • 레이저 용접 분야 : FLOW-3D는 용융 풀 역학분야에서  용접 이음매의 일관성과 강도에 중요한 역할을 하는 레이저 용접에 사용됩니다.
  • 자동차 산업 분야 :  FLOW-3D는 변속기 및 기어 모델링, 냉각 및 연료 시스템에서 가장 까다로운 열 유동 해석 분야의 문제 해결에 많이 사용되고 있습니다.  
  • 소비자 제품 개발 분야 : FLOW-3D는 샴푸 및 세제 병 채우기와 같은 소비자 제품 개발에 널리 사용됩니다 .
  • 미세 유체 분야 : FLOW-3D는 현재 3D 프린팅으로 제품을 개발하고 있는 잉크젯 프린팅 산업에서 연구 개발부문에 압도적으로 사용되고 있습니다. 
  • 생물 의학 응용 분야 : 자유 표면 분석 필요성이 발생할 때마다,  FLOW-3D 는 이러한 시스템을 시뮬레이션하는데 탁월합니다.
  • 코팅분야 연구개발 :  FLOW-3D는 코팅분야 제품 연구개발에 매우 자주 사용되고 있습니다.

회사 소개

소개

company_banner
  • 설립일 : 1997 년 11 월
  • 위치 : ( 본사 ) 서울시 금천구 가산동 우림라이온스밸리 B 동 301~2 호

사업 영역

당사는 엔지니어링 전문 기술서비스 업체로 다음과 같은 분야에서 탁월한 기술 서비스를 제공하고 있습니다 .

  • 국내유일의 수치해석 프로그램 (FLOW-3D) 공급 및 기술지원 – 한국총판
  • 수치해석 기술 컨설팅 전문 회사로 오랜 경험과 노하우를 바탕으로 차별화된 전문 기술 서비스 제공
  • 해양, 댐, 수처리, 주조, 철강, 조선, 항공/우주 분야, MEMS, 바이오 등 다양한 산업분야의 전문 해석
  • 용접 , Laser Welding, 3D 프린팅 해석
  • 첨단 제조 , 제철 , 우주항공 , 바이오 분야 수치해석컨설팅 및 연구용역
  • 엔지니어링 소프트웨어 개발
  • 사이펀 여수로 , 취수사이펀 시공설치

대표이사 인사말

(주)에스티아이C&D는 고품질 엔지니어링 서비스인 수치해석, 엔지니어링 S/W 개발 등 선진 기술개발에 힘쓰고 있습니다.

국내 최고의 기술력과 수치해석 경험을 통해 수처리분야의 댐해석, 강과 하천, 정수장, 하폐수처리장, 바이오 배양시설 및 첨단 제품 개발을 위한 자동차 부품개발, 전자, 주조, 철강 등 고난이도 수치해석 컨설팅 및 엔지니어링 소프트웨어를 개발 제공하고 있습니다.

고객의 성공이 최상의 가치임을 알기에 언제나 고객 여러분의 의견에 귀 기울이고 고객과 함께 성장하도록 최선을 다하겠습니다. 여러분의 관심과 조언을 부탁 드립니다.

대표이사 홍기원

주요 고객사

STI C&D 주요 고객사
STI C&D 주요 고객사

수치해석 (CFD) 분야 소개

FLOW-3D 공급 기술지원

  • FLOW-3D 의 국내 독점 공급 및 기술지원
  • FLOW-3D 의 정기 및 비정기 교육지원과 기술지원
  • 국내 대학 및 국가연구기관 교육 및 세미나

엔지니어링 수치해석 기술 컨설팅 용역

  • 홍수 시 하천 , 수리구조물의 3 차원적 영향검토
  • 댐 , 여수로 , 갑문 등 수리구조물의 다양한 설계안 3 차원 동수역학적 검토 및 최적 설계안 도출
  • 정수장 , 하수처리장 3 차원 유량분배 , 유동현상 등의 검토
  • 발전소 취배수로의 3 차원 유동해석을 통한 평면배치 및 적정규모 검토 , 증설 영향등 검토
  • 기계 , MEMS, 항공 , 조선 등 다양한 분야의 열전달 및 유체 유동에 대한 해석
  • 금형설계 및 주조품 결함예측에 관련된 Die Casting 분야 해석
  • 제강 , 연주 공정 , 열유동 , 응고해석 등 철강분야 3 차원 유동해석

오시는 길