**2. Microfluidic technology**

#### **2.1. What microfluidic has to offer?**

Microfabrication enables the deposition and etching of thin layers (angstrom to micrometer) of different materials on silicon or glass substrates. These layers can be patterned with accuracy and high resolution, down to the nanometer level using lithography. Lithography is the technique used to transfer a pattern from a mask to the substrate to control the location of the deposition of the next layer or the etching of an existing layer on the substrate.

Microfabricated devices have been used in a broad range of biomedical and biological applications. In the last decade, microchips have been used in microscale sensors for surgical instruments, to monitor physiological activities, in microfluidics applications such as drug discovery and delivery, cell sorting, DNA amplification, electrophoresis, etc. Of relevance to this chapter are the microchips for blood fractionation and cell sorting.

Micromachining consists of a series of robust, well controlled, and well characterized processes that enable the fabrication of microfluidic devices. Such devices can be made costeffective by the use of any of a wide range of biocompatible polymers or plastics and bulk processing (mass production). The miniaturization of reactions and assays confers many advantages over "macro" scale techniques beyond the obvious reduction in quantities of reagents and materials required per test. The scaling down of volumes results in higher surface to volume ratio: a cube with side length L will scale down by a factor of L3, while the surface area will scale down by only a factor of L2. Thus, miniaturization results in higher reactivities, shorter diffusion distances, smaller heat capacities, faster heat exchange, shorter assay times, and better overall process control, as well as the capability to integrate multiple steps and to achieve massive parallelization on-chip. Additionally, microfluidic devices are safer than "macro" platforms due to the smaller chemical quantities used and hence the lower stored energies. A microfluidic device that performs one assay is typically referred to as Lab-on-a-chip, while a device that integrates more than one step is referred to as micro total analysis system (µTAS) [18].

#### **2.2. Introduction to microfabrication**

198 Blood Cell – An Overview of Studies in Hematology

separation and analyses.

substrate.

**2. Microfluidic technology** 

**2.1. What microfluidic has to offer?** 

marker/antibody to exclusively detect CTCs or fetal cells [32].

The number of available cells of interest poses and additional challenge; in some applications such as the isolation of CTCs from cancer patient or fetal nucleated red blood from maternal circulation, only 1-2 cell are available per milliliter (mL) of whole blood. Specificity is problematic for either method, due to the lack of a cell-specific surface

Despite great successes, cell sorting techniques are not ideal and therefore remain an active area of research. In addition to sensitivity and specificity requirements, an ideal technique should not be overly labor intensive, should be automated and quantitative, the results should predict clinical outcome, and help the physician personalize therapeutic options. Automating sample preparation and handling would minimize human errors. Integration of preparation, cell sorting, and post processing will lead to more cost-effective instruments, and alleviate the need for trained personnel and infrastructure. Microfluidic technology enables the precise control over the cell microenvironment during separation, scales down the analyses to very small volume of blood, and has the potential for high-throughput to cell

Microfabrication enables the deposition and etching of thin layers (angstrom to micrometer) of different materials on silicon or glass substrates. These layers can be patterned with accuracy and high resolution, down to the nanometer level using lithography. Lithography is the technique used to transfer a pattern from a mask to the substrate to control the location of the deposition of the next layer or the etching of an existing layer on the

Microfabricated devices have been used in a broad range of biomedical and biological applications. In the last decade, microchips have been used in microscale sensors for surgical instruments, to monitor physiological activities, in microfluidics applications such as drug discovery and delivery, cell sorting, DNA amplification, electrophoresis, etc. Of

Micromachining consists of a series of robust, well controlled, and well characterized processes that enable the fabrication of microfluidic devices. Such devices can be made costeffective by the use of any of a wide range of biocompatible polymers or plastics and bulk processing (mass production). The miniaturization of reactions and assays confers many advantages over "macro" scale techniques beyond the obvious reduction in quantities of reagents and materials required per test. The scaling down of volumes results in higher surface to volume ratio: a cube with side length L will scale down by a factor of L3, while the surface area will scale down by only a factor of L2. Thus, miniaturization results in higher reactivities, shorter diffusion distances, smaller heat capacities, faster heat exchange, shorter assay times, and better overall process control, as well as the capability to integrate multiple steps and to achieve massive parallelization on-chip. Additionally, microfluidic devices are

relevance to this chapter are the microchips for blood fractionation and cell sorting.

This section will briefly describe the basic concepts in the microfabrication of microfluidic devices. Microfabrication is the already subject of many textbooks and the interested reader can consult one of these for more in-depth details [33, 34]. Microfabrication is the technology developed by ICs industry to make devices and circuits with feature sizes as small as 14 nm in research, and 45 nm in production. Among these are the microprocessors and the electronic components found in computers, smart phones, television sets, and major other electronic products. Microfabrication is also used for MEMS devices (micro-electro mechanical systems), devices that include a movable part and can be used for sensing. The airbag sensor used to deploy an airbag in a vehicle, the pressure sensors inside car tires, and the electronic compass in a smart phone are all examples of MEMS devices that we unknowingly use every day. BioMEMS, microfluidic devices, and µTAS are all microfabricated devices similar to MEMS but customized for biological and chemical applications.

Creation of a a new microfluidic device includes design of the channel(s), fluid inlet(s) and outlet(s), using a CAD (computer-aided design) software. These softwares, such as CoventorWare® [35], ANSYS CFD [36], COMSOL Multiphysics® [37], can be also used for simulation of the various design parameters such as device dimensions, heat transfer, and flow conditions, therefore narrowing the design space range in which optimum performance should be obtained. The pattern of channels is laid out with the CAD software; this is the 2D design of the device. The depth of the channel will be determined by the etching time. The drawings are transferred to a mask, typically a glass or quartz plate (transparent to UV light), covered with chrome. The chrome is etched (removed) where the UV light will expose the photoresist. The mask, similar to a stencil, transfers the pattern to the photoresist. The device is built on a substrate, which is a silicon, glass or quartz wafer, or a regular glass slide. After substrate cleaning, the photoresist, a photosensitive polymer, is applied. Photoresist is dispensed onto the substrate and it is spun at high speed (2000- 4000rpm) to create a thin (1-100 micrometer), uniform and smooth layer. The mask is placed in contact with the substrate and exposed to UV lights on a mask aligner. The photoresist is developed in a developer solution specific for it, and is removed from areas exposed to UV light (positive photoresist). The channels can be etched, with substrate material being removed from areas unprotected by the photoresist. Etching can be either wet (using chemicals) or dry (plasma etching). Deep reactive ion etching (DRIE) is a plasma etching technique typically used to achieve deep channels with vertical side walls. Use of DRIE is necessary if the etched substrate will serve as a template for molding devices in polymer such that the polymer mold can be peeled off the substrate. At this stage, the device can be

sealed by the bonding of a top piece to the substrate. Top pieces are typically transparent, to permit observation under the microscope. Silicon can be bonded to glass by thermal or anodic bonding technique. Usually, however, microfluidic devices are constructed of polymer. Polymers are more cost effective, transparent, and biocompatible materials, many polymer devices can be molded from one silicon master, this is a key advantage because microfluidics devices are hard to clean and hence, can only be used for one or a few experiments.

Use of Microfluidic Technology for Cell Separation 201

**2.3. Types of flow: laminar versus turbulent** 

(Re), which is the ratio of inertial forces to the viscous forces:

V is the mean velocity of the object relative to the fluid (SI units m/s), L is a characteristic linear dimension, (travelled length of the fluid, and µ is the dynamic viscosity of the fluid (Pa·s or N·s/m² or kg/(m·s)).

1 to 100) thus only laminar flow is relevant for microfluidic devices.

Where:

is the density of the fluid (kg/m³),

**3. Fluid transport process** 

pressure, and radius of the tube:

**3.1. Poiseuille flow** 

Two modes of fluid flow exist: laminar and turbulent. In laminar flow, the fluid moves with slow velocity and each particle of fluid moves in a straight trajectory parallel to the channel walls in the flow direction; and the velocity, pressure, and other flow properties at each point in the fluid remain constant. There are no cross-currents perpendicular to the flow direction, no eddies or swirls of current in laminar flow. Examples are oil flowing slowly through a tube, and blood flowing through a capillary. Turbulent flow in contrast is chaotic, with rapid, spatial and temporal variations of pressure and velocity. Examples of turbulent flow are the blood flow through in arteries, the flow of water through pumps and turbines,

and the flow eddying seen in boat wakes and around the wing tips of aircraft [17, 33].

*e <sup>L</sup> R v*

The relative turbulence of a flow can be determined by the dimensionless Reynolds number

Below a certain Re value the flow is laminar; above this threshold the flow becomes turbulent. For macroscopic structures such as pipes with a circular cross-section, the transition from laminar to turbulent flow has been empirically determined at Re of approximately 2300. For most microstructures, in contrast, the Re number is usually low (10-

A pressure-driven laminar flow inside a tube with a circular cross-section, away from the entrance, is commonly known as *Hagen-Poiseuille* flow or simply *Poiseuille* flow [17, 33]. This flow is governed by the *Navier-Stokes* equations, which are nonlinear, second-order, partial differential equations for describing incompressible fluids. These equations are derived from motion and conservation of mass equations. *Poiseuille* flow is a solution to the *Navier-Stokes* equation, and describes the fluid velocity at any point as a function of the viscosity,

<sup>1</sup> 2 2

*<sup>p</sup> V rR z* 

4 *<sup>z</sup>*

Polydimethysiloxane (PDMS), PMMA, polyurethane, and polystyrene are all polymeric materials used for microfluidic devices. PDMS, the most frequently used, comes in the form of two liquid components that are mixed (1:10, w/w) and poured onto the substrate. The PDMS is degassed to remove all air bubbles and ensure that the liquid fills the smallest feature of the mold. PDMS is cured in a 60-80oC oven for 20 to 45 min to solidify. Once solid, it can be peeled off the master substrate. PDMS devices are sealed with glass cover slips to form the channels. Several alternatives to the above described processes exist that can accelerate fabrication. Masks can be printed on transparencies using high-resolution printers; this method is suitable for feature sizes of 100 µm or larger. SU8 is a photoresist that is used to create deep structures for molding eliminating the need for deep etching such as the utilization of DRIE. Utilization of a chrome mask and a well-equipped clean-room facility are necessary for making devices with very small features; sub-micrometer to a few micrometers. The development cycle from concept to prototype can take few weeks. For less fine featured devices, the use of transparencies, polymers, and/or SU8 can be reduce the cost and developmental time to one to few days from concept to prototype [38]. Figure 1a and 1b illustrates the typical processes involved in making channels in silicon and molding devices in PDMS respectively.

**Figure 1.** a: Schematic summary of the processes involved in making channel in a hard substaret such as silicon. b: Schematic summary of steps involved in making a PDMS mold from a hard master.

#### **2.3. Types of flow: laminar versus turbulent**

Two modes of fluid flow exist: laminar and turbulent. In laminar flow, the fluid moves with slow velocity and each particle of fluid moves in a straight trajectory parallel to the channel walls in the flow direction; and the velocity, pressure, and other flow properties at each point in the fluid remain constant. There are no cross-currents perpendicular to the flow direction, no eddies or swirls of current in laminar flow. Examples are oil flowing slowly through a tube, and blood flowing through a capillary. Turbulent flow in contrast is chaotic, with rapid, spatial and temporal variations of pressure and velocity. Examples of turbulent flow are the blood flow through in arteries, the flow of water through pumps and turbines, and the flow eddying seen in boat wakes and around the wing tips of aircraft [17, 33].

The relative turbulence of a flow can be determined by the dimensionless Reynolds number (Re), which is the ratio of inertial forces to the viscous forces:

$$\mathcal{R}\_e = \rho v \frac{L}{\mu}$$

Where:

200 Blood Cell – An Overview of Studies in Hematology

experiments.

in PDMS respectively.

sealed by the bonding of a top piece to the substrate. Top pieces are typically transparent, to permit observation under the microscope. Silicon can be bonded to glass by thermal or anodic bonding technique. Usually, however, microfluidic devices are constructed of polymer. Polymers are more cost effective, transparent, and biocompatible materials, many polymer devices can be molded from one silicon master, this is a key advantage because microfluidics devices are hard to clean and hence, can only be used for one or a few

Polydimethysiloxane (PDMS), PMMA, polyurethane, and polystyrene are all polymeric materials used for microfluidic devices. PDMS, the most frequently used, comes in the form of two liquid components that are mixed (1:10, w/w) and poured onto the substrate. The PDMS is degassed to remove all air bubbles and ensure that the liquid fills the smallest feature of the mold. PDMS is cured in a 60-80oC oven for 20 to 45 min to solidify. Once solid, it can be peeled off the master substrate. PDMS devices are sealed with glass cover slips to form the channels. Several alternatives to the above described processes exist that can accelerate fabrication. Masks can be printed on transparencies using high-resolution printers; this method is suitable for feature sizes of 100 µm or larger. SU8 is a photoresist that is used to create deep structures for molding eliminating the need for deep etching such as the utilization of DRIE. Utilization of a chrome mask and a well-equipped clean-room facility are necessary for making devices with very small features; sub-micrometer to a few micrometers. The development cycle from concept to prototype can take few weeks. For less fine featured devices, the use of transparencies, polymers, and/or SU8 can be reduce the cost and developmental time to one to few days from concept to prototype [38]. Figure 1a and 1b illustrates the typical processes involved in making channels in silicon and molding devices

**Figure 1.** a: Schematic summary of the processes involved in making channel in a hard substaret such as silicon. b: Schematic summary of steps involved in making a PDMS mold from a hard master.

 is the density of the fluid (kg/m³), V is the mean velocity of the object relative to the fluid (SI units m/s), L is a characteristic linear dimension, (travelled length of the fluid, and µ is the dynamic viscosity of the fluid (Pa·s or N·s/m² or kg/(m·s)).

Below a certain Re value the flow is laminar; above this threshold the flow becomes turbulent. For macroscopic structures such as pipes with a circular cross-section, the transition from laminar to turbulent flow has been empirically determined at Re of approximately 2300. For most microstructures, in contrast, the Re number is usually low (10- 1 to 100) thus only laminar flow is relevant for microfluidic devices.

#### **3. Fluid transport process**

#### **3.1. Poiseuille flow**

A pressure-driven laminar flow inside a tube with a circular cross-section, away from the entrance, is commonly known as *Hagen-Poiseuille* flow or simply *Poiseuille* flow [17, 33]. This flow is governed by the *Navier-Stokes* equations, which are nonlinear, second-order, partial differential equations for describing incompressible fluids. These equations are derived from motion and conservation of mass equations. *Poiseuille* flow is a solution to the *Navier-Stokes* equation, and describes the fluid velocity at any point as a function of the viscosity, pressure, and radius of the tube:

$$V\_z = \frac{1}{4\,\mu} \left(\frac{\partial p}{\partial z}\right) \left(r^2 - R^2\right),$$

Where:

Vz is the velocity distribution, µ is the fluid viscosity, (p/z) is the z component of the pressure gradient, R is the distance from the center of the tube, and R is the radius of the tube.

This equation reveals that the velocity distribution is parabolic, with the maximum velocity occurring at the center of the tube. **Figure 2a**

Use of Microfluidic Technology for Cell Separation 203

This equation reveals that the velocity is linearly proportional to the applied electric field, such that all travel at the same speed inside the channel **Figure 2b**. This situation contrasts with the pressure-driven flow, in which the middle has the greatest velocity resulting, in a

In conclusion, the velocity distribution being parabolic or flat can impact the microfluidic device performance. The electrokinetic flow is favored in applications such as DNA separation, proteomics, rapid mixing, and time dependent applications where sample homogeneity and reaction time need to be stringently controlled. The flow pattern is typically not an issue for cell separation since it happens in continuous flow without time

**Figure 2.** a: schematic of velocity distribution in Poiseuille flow. b: schematic of velocity distribution in Electro-osmosis flow.

**4. Technologies and on-chip mechanisms of separation** 

separation is enrichment rather than a true purified population [9].

Cells are separated either in bulk or individually. In individual cell sorting, each cell is analyzed, and then the cells of choice are individually selected. This technique is rarely used, due to its very low throughput. Thus cells are generally sorted by bulk separation, in which a large number of cells are selected on the basis of shared characteristics such size, density, or the affinity of a receptor for a specific cell-surface target. The result of such bulk

The cells of interest are first identified, than separated, and finally collected. The initial step is to screen the cells of interest to identify one or more common specific characteristics to be used as the basis of sorting. Specific characteristics can be intrinsic such as size, density, response to electrical or magnetic fields, or resistance to chemical lysis. Alternatively, cells can be labeled using a specific cell surface target that binds to a monoclonal antibody

Once the cells of interest have been identified, they can be separated from the other cell types. In some methods, the identification and separation occur simultaneously, e.g., affinity

conjugated to a fluorophore, or to magnetic particles for flow cytometric cell sorting.

µ is the Viscosity.

parabolic migration.

dependence.

capture.

#### **3.2. Electrokinetic Flow**

Electrokinetic flow is the underlying basis of electro-osmosis, electrophoresis, streaming potential, and dielectrophoresis. In electro-osmosis, fluid is made to flow through, by the application of an electric field. The field induces the formation of an electric double layer (EDL). This EDL consists of (1) a compact liquid layer adjacent to the channel surface that has immobile balancing charges, and (2) a second *diffuse* liquid layer, composed of mobile ions. Most solid-liquid, and many liquid-liquid, interfaces have an electrostatic charge, and hence an electric field exists there. In the presence of an electric field, molecules of many dielectric materials, such as the glass or polymer used for making the microfluidic devices, will become permanently polarized, since the material has a dipole that comprises two opposite, but equal, charges, due to the asymmetrical molecular structure. The electrostatic charges on the channel surface, mainly negative charges in the case of glass or polymer, attract counter-ions from the liquid. This attraction creates the channel double layer; the first has immobile charges that balance the charges on the channel surface. The second *diffuse* layer has a higher concentration of counter-ions near the channel surface than does the bulk of the fluid. The net charge density gradually decreases to zero in the bulk liquid. The diffuse layer will move under the electric field. Surrounding molecules are pulled along by a viscous effect, resulting in bulk fluid motion or electro-osmotic flow. The diffuse layer is several nm to 1 or even 2 µm thick, depending on the ionic concentration and electrical properties of the liquid. The *Poisson-Boltzmann* equation describes the ion and potential distribution in the diffuse layer. The electro-osmotic flow velocity can be quantified by Li's equation [39]:

$$V\_{\omega v} = E\_z \varepsilon\_r \varepsilon\_o \frac{\tilde{\zeta}}{\mu}$$

Where,

Vav is the average elelctro-osmotic flow velocity, z is the applied electric field (V/m), r is the Dielectric constant of the medium, o is the permetivity of the vacuum, is the zeta potential at the shear plane, and

µ is the Viscosity.

202 Blood Cell – An Overview of Studies in Hematology

(p/z) is the z component of the pressure gradient, R is the distance from the center of the tube, and

occurring at the center of the tube. **Figure 2a**

This equation reveals that the velocity distribution is parabolic, with the maximum velocity

Electrokinetic flow is the underlying basis of electro-osmosis, electrophoresis, streaming potential, and dielectrophoresis. In electro-osmosis, fluid is made to flow through, by the application of an electric field. The field induces the formation of an electric double layer (EDL). This EDL consists of (1) a compact liquid layer adjacent to the channel surface that has immobile balancing charges, and (2) a second *diffuse* liquid layer, composed of mobile ions. Most solid-liquid, and many liquid-liquid, interfaces have an electrostatic charge, and hence an electric field exists there. In the presence of an electric field, molecules of many dielectric materials, such as the glass or polymer used for making the microfluidic devices, will become permanently polarized, since the material has a dipole that comprises two opposite, but equal, charges, due to the asymmetrical molecular structure. The electrostatic charges on the channel surface, mainly negative charges in the case of glass or polymer, attract counter-ions from the liquid. This attraction creates the channel double layer; the first has immobile charges that balance the charges on the channel surface. The second *diffuse* layer has a higher concentration of counter-ions near the channel surface than does the bulk of the fluid. The net charge density gradually decreases to zero in the bulk liquid. The diffuse layer will move under the electric field. Surrounding molecules are pulled along by a viscous effect, resulting in bulk fluid motion or electro-osmotic flow. The diffuse layer is several nm to 1 or even 2 µm thick, depending on the ionic concentration and electrical properties of the liquid. The *Poisson-Boltzmann* equation describes the ion and potential distribution in the diffuse layer. The electro-osmotic flow velocity can be quantified by Li's

*V E av z r o*

 

Vz is the velocity distribution, µ is the fluid viscosity,

R is the radius of the tube.

**3.2. Electrokinetic Flow** 

equation [39]:

Vav is the average elelctro-osmotic flow velocity,

is the zeta potential at the shear plane, and

z is the applied electric field (V/m), r is the Dielectric constant of the medium, o is the permetivity of the vacuum,

Where,

Where:

This equation reveals that the velocity is linearly proportional to the applied electric field, such that all travel at the same speed inside the channel **Figure 2b**. This situation contrasts with the pressure-driven flow, in which the middle has the greatest velocity resulting, in a parabolic migration.

In conclusion, the velocity distribution being parabolic or flat can impact the microfluidic device performance. The electrokinetic flow is favored in applications such as DNA separation, proteomics, rapid mixing, and time dependent applications where sample homogeneity and reaction time need to be stringently controlled. The flow pattern is typically not an issue for cell separation since it happens in continuous flow without time dependence.

**Figure 2.** a: schematic of velocity distribution in Poiseuille flow. b: schematic of velocity distribution in Electro-osmosis flow.
