**MEMS-Based Microdevice for Cell Lysis and DNA Extraction**

Xing Chen, Dafu Cui, Haoyuan Cai, Hui Li, Jianhai Sun and Lulu Zhang *State Key Lab. of Transducer Tech., Inst. of Electronics, Chinese Academy of Sciences China* 

#### **1. Introduction**

22 Microelectromechanical Systems and Devices

Zrenner, E.; Stett, A.; Weiss, S.; Aramant, R.B.; Guenther, E.; Kohler, K.; Miliczek, K.-D.;

Zolfaghari, A. Chan, A. & Razavi, B., Stacked Inductors and Transformers in CMOS

2567, ISSN 0042-6989.

0018-9200.

Seiler, M.J. & Haemmerle, H. (1999). Can Subretinal Microphotodiodes Successfully Replace Degenerated Photoreceptors? *Vision Research*, Vol. 39, (July 1999), pp. 2555-

Technology. *IEEE J. of Solid-State Circuits*, Vol. 36, (April 2001), pp. 620-628, ISSN

With the development of microelectromechanical system (MEMS) technology, micrototal analytical systems (μTAS) which has the potential for integrating sample pretreatment, target amplification, and detection, has been in progress. Micromachined analytical systems have several advantages over their large-scale counterparts, including low cost, disposability, low reagent and sample consumption, portability, and lower consumption. Many such devices have been demonstrated in the literature, including PCR microchips (Northrup et al., 1993; Copp et al., 1998; Panaro et al., 2005), DNA microchips (Fan et al., 1999), DNA biosensors (Kwakye et al., 2006), capillary electrophoresis (CE) microchips (Harrison et al., 1993; Backhouse et al., 2003; Liu et al., 2006), protein microchips (Yang et al., 2001; Wilson & Nie, 2006), etc. Most of these analytical processes need an effective yet simple method of obtaining high-quality DNA. Hence miniature devices for rapid sample pretreatment of DNA, including cell lysis and genomic DNA purification, are crucial for genetic application.

Traditional phenol extraction is a complex and time-consuming method for extracting DNA, and even some commercial purification kits require several centrifugal operations. The implementation of DNA purification on a microdevice is initially demonstrated based on the principle of solid phase extraction (SPE). The SPE on-microdevice can minimize sample loss and contamination problems as well as reduce analysis time, and besides, this SPE method can avoid problems of physical and biochemical degradation of DNA. For example, Tian et al. (Tian et al., 2000) established an SPE DNA purification microdevice in a capillary packing with silica resin matrix which could extract enough DNA for PCR reaction. Wolfe et al. (Wolfe et al., 2002) and Breadmore et al. (Breadmore et al., 2003) immobilized bare-silica beads matrix in microchannels by sol–gel technology for DNA purification. But a high packing density for larger surface area in the microfluidic device results in problems of backpressure and clogging of crude samples, and what is more, it is difficult to control the small particles in microdevices. A micropillar array fabricated by MEMS technology in a microchamber or channel increases the surface area available for DNA adsorption (Christel et al., 1999; Cady et al., 2003). However, the increasing surface area is limited and the problems of clogging could not be completely solved. Hence, a novel solid-phase matrix which should be easily integrated in microdevices is under demand.

MEMS-Based Microdevice for Cell Lysis and DNA Extraction 25

electrochemical etching technology, and then were characterized by using SEM (scanning electron microscopy) and BET (Brunauer, Emmet, and Teller) nitrogen adsorption technique. The porous layer directly generated on the internal walls of the channels can greatly enhance the "active" surface area compared to the non-anodized one, and also avoid

In this integrated microdevice, cells were rapidly lyzed at first, and then genomic DNA was released and absorbed on a solid-phase matrix. Secondly, washing buffer was pumped through this microdevice for removing the proteins and other impurities which were also absorbed during the previous process. Finally, elution buffer was used to elute DNA and the desorbed DNA was collected at the outlet port. During continuous flow process, on-line rapid cell lysis and PCR amplifiable genomic DNA purification on the single MEMS-based microdevice have been implemented, which confirmed that the proposed microfluidic device was capable of directly providing genic analyte for the following molecular biology

In microdevice channels, the flow is laminar in nature. Reynolds number, which is the ratio of the inertial forces to the viscous forces, can be used to characterize the flow. It is given by

Where ρ is the fluid density, ν is the flow velocity, and μ is the fluid viscosity. Dh represents

 Dh = 2 a b/(a+b) (2) Where a and b are the cross sectional dimensions of the channel. The Reynolds numbers involved in the field of microfluidic device are usually below 100 (Gravesen et al., 1993),

According the formula (1), the fluid density (ρ) and the fluid viscosity (μ) are constants. Re is depended on the flow velocity (ν) and the hydraylic diameter (Dh). For our microdevice with a 200µm wide and 100µm deep rectangle channel, the hydraylic diameter (Dh) calculated by the formula (2) is 133µm. Note that the blood density (ρ) is 1.05g/cm3, and the viscosity of blood (μ) is 3.115mPa.s. According to the formula (1), the Re in our experimental was 0.035~8.4, when the flow velocity (ν) is about 0.0008~0.024m/s. Therefore a typical laminar flow is found in our experimental. Lack of turbulent flow limits the effective mixing of the fluids under investigation. On the other hand, for the cell lysis on the microdevice based on the chemical method, it requires that each cell should be fully exposed to lysis reagents in microchannels, which is determined by mixing performance of the microdevice. In microdevice channels where flow is laminar in nature, the mixing is dictated by diffusion. Average time for a small portion of a fluid to diffuse a distance L can be estimated by

 TD = L2 / D (3) Where D is the diffusion coefficient of the liquid. The above equation can be used to predict the order of time scale of mass diffusion. As it suggests, one can dramatically reduce the mixing time by reducing the diffusion length required for mixing or increasing the contact

area between two different liquids while keeping the volume constant.

the hydraylic diameter and, for rectangular channels, is given by

Re = ρ ν Dh / μ (1)

the problems of packing solid phase matrix.

research or medical assay.

which prevent turbulent flow.

**2. Theory** 

It is well known that porous silicon with a relatively large specific surface area-to-volume (hundreds of square meters per cubic centimeter) can significantly increase the available interface area. Porous silicon can be easily obtained by electrochemical etching technology, and the pore geometry, the surface morphology, and the porosity of porous silicon can be very precisely controlled by the electrochemical etching conditions (Bisi et al., 2000; Schmuki et al., 2003). Porous silicon with several useful characteristics has already found applications in diverse fields such as solar cells (Bilyalov et al., 2003), RF (Park et al., 2001), (bio)chemical sensors (Massera et al., 2004; Björkqvist et al., 2004), etc.. Furthermore porous silicon with huge surface area as a kind of biocompatible solid support has been used to absorb enzyme, protein and other biologic molecules in fields of the enzyme micro reactors (Melander et al., 2005; Bengtsson et al., 2002), chromatography (Clicq et al., 2004), and antibody micro arrays (Steinhauer et al., 2005). Due to the fact that the technology for porous silicon fabrication is compatible with standard microelectronic and MEMS techniques, we have made use of porous silicon as the solid-phase matrix to extract DNA (Chen et al., 2006, 2007).

DNA purification from initial samples requires disrupting cells to liberate the nucleic acids before SPE process. Typical laboratory protocols for lysis steps include the use of enzymes (lysozyme), chemical lytic agents (detergents), and mechanical forces (sonication, bead milling). However, many such lysis techniques are not amenable to be implemented in a microfluidic platform. Miniaturization methods for lyzing cells are required so that cell lysis can be integrated into the μTAS. Therefore, a desirable lysis method should rapidly destroy cells while at the same time, it could not destroy nucleic acids and inhibit the following PCR reaction either, which would be amenable to be integrated with SPE on microdevices. Carlo et al. (Carlo et al., 2003) reported a mechanical cell lysis microdevice with nanostructural barbs which was used to disrupt sheep blood cells. However, the fabrication process of nanostructures is complex. Other miniaturization cell lysis methods include thermal (Lee & Tai, 1999), electrical (Gao et al., 2004; Taylor et al., 2001), ultrasonication (Belgrader et al., 2000; LaMontagne1 et al., 2002), and chemical treatments (Li & Jed Harrison, 1997). These approaches have been shown to be moderately successful. However, they all depend on the use of an external power supply and the devices may be quite complicated and costly to fabricate. Hence, chemical disruption methods are chosen due to the fact that they can be compatible with SPE on microdevices and do not require complex process of fabrication either. For example, Schilling et al. (Schilling et al., 2002) reported a relatively simple T-type microfluidic device that allowed the continuous lysis of bacterial cells using β-galactosidase. Sethu et al. (Sethu et al., 2004) established a continuous-flow microfluidic device for rapid erythrocyte lysis.

In this chapter, we have described a novel integrated MEMS-based microdevice capable of performing online cell lysis and genomic DNA purification during continuous flowing process. The method of chemical disruption was chosen for cell lysis, while the method of solid phase extraction was chosen for DNA purification. For cell lysis microdevice based on chemical disruption method, mixing is the key step due to the fact that each cell should be fully exposed to lysis buffer before the total lysis of cells, which is limited by the mixing speed. The mixing procedure was numerically simulated. Based on the results of the simulation, the mixing model for cell lysis was optimized to construct the microfluidic devices fabricated by MEMS technology. For DNA purification microdevice based on the SPE principle, solid-phase matrix is the key factor. The porous silicon was used as the solidphase matrix to extract DNA. The porous silicon dioxide matrix was fabricated by electrochemical etching technology, and then were characterized by using SEM (scanning electron microscopy) and BET (Brunauer, Emmet, and Teller) nitrogen adsorption technique. The porous layer directly generated on the internal walls of the channels can greatly enhance the "active" surface area compared to the non-anodized one, and also avoid the problems of packing solid phase matrix.

In this integrated microdevice, cells were rapidly lyzed at first, and then genomic DNA was released and absorbed on a solid-phase matrix. Secondly, washing buffer was pumped through this microdevice for removing the proteins and other impurities which were also absorbed during the previous process. Finally, elution buffer was used to elute DNA and the desorbed DNA was collected at the outlet port. During continuous flow process, on-line rapid cell lysis and PCR amplifiable genomic DNA purification on the single MEMS-based microdevice have been implemented, which confirmed that the proposed microfluidic device was capable of directly providing genic analyte for the following molecular biology research or medical assay.

#### **2. Theory**

24 Microelectromechanical Systems and Devices

It is well known that porous silicon with a relatively large specific surface area-to-volume (hundreds of square meters per cubic centimeter) can significantly increase the available interface area. Porous silicon can be easily obtained by electrochemical etching technology, and the pore geometry, the surface morphology, and the porosity of porous silicon can be very precisely controlled by the electrochemical etching conditions (Bisi et al., 2000; Schmuki et al., 2003). Porous silicon with several useful characteristics has already found applications in diverse fields such as solar cells (Bilyalov et al., 2003), RF (Park et al., 2001), (bio)chemical sensors (Massera et al., 2004; Björkqvist et al., 2004), etc.. Furthermore porous silicon with huge surface area as a kind of biocompatible solid support has been used to absorb enzyme, protein and other biologic molecules in fields of the enzyme micro reactors (Melander et al., 2005; Bengtsson et al., 2002), chromatography (Clicq et al., 2004), and antibody micro arrays (Steinhauer et al., 2005). Due to the fact that the technology for porous silicon fabrication is compatible with standard microelectronic and MEMS techniques, we have made use of

porous silicon as the solid-phase matrix to extract DNA (Chen et al., 2006, 2007).

erythrocyte lysis.

DNA purification from initial samples requires disrupting cells to liberate the nucleic acids before SPE process. Typical laboratory protocols for lysis steps include the use of enzymes (lysozyme), chemical lytic agents (detergents), and mechanical forces (sonication, bead milling). However, many such lysis techniques are not amenable to be implemented in a microfluidic platform. Miniaturization methods for lyzing cells are required so that cell lysis can be integrated into the μTAS. Therefore, a desirable lysis method should rapidly destroy cells while at the same time, it could not destroy nucleic acids and inhibit the following PCR reaction either, which would be amenable to be integrated with SPE on microdevices. Carlo et al. (Carlo et al., 2003) reported a mechanical cell lysis microdevice with nanostructural barbs which was used to disrupt sheep blood cells. However, the fabrication process of nanostructures is complex. Other miniaturization cell lysis methods include thermal (Lee & Tai, 1999), electrical (Gao et al., 2004; Taylor et al., 2001), ultrasonication (Belgrader et al., 2000; LaMontagne1 et al., 2002), and chemical treatments (Li & Jed Harrison, 1997). These approaches have been shown to be moderately successful. However, they all depend on the use of an external power supply and the devices may be quite complicated and costly to fabricate. Hence, chemical disruption methods are chosen due to the fact that they can be compatible with SPE on microdevices and do not require complex process of fabrication either. For example, Schilling et al. (Schilling et al., 2002) reported a relatively simple T-type microfluidic device that allowed the continuous lysis of bacterial cells using β-galactosidase. Sethu et al. (Sethu et al., 2004) established a continuous-flow microfluidic device for rapid

In this chapter, we have described a novel integrated MEMS-based microdevice capable of performing online cell lysis and genomic DNA purification during continuous flowing process. The method of chemical disruption was chosen for cell lysis, while the method of solid phase extraction was chosen for DNA purification. For cell lysis microdevice based on chemical disruption method, mixing is the key step due to the fact that each cell should be fully exposed to lysis buffer before the total lysis of cells, which is limited by the mixing speed. The mixing procedure was numerically simulated. Based on the results of the simulation, the mixing model for cell lysis was optimized to construct the microfluidic devices fabricated by MEMS technology. For DNA purification microdevice based on the SPE principle, solid-phase matrix is the key factor. The porous silicon was used as the solidphase matrix to extract DNA. The porous silicon dioxide matrix was fabricated by In microdevice channels, the flow is laminar in nature. Reynolds number, which is the ratio of the inertial forces to the viscous forces, can be used to characterize the flow. It is given by

$$\text{Re} = \text{p} \,\text{v} \,\text{D}\_{\text{h}} / \,\text{p} \tag{1}$$

Where ρ is the fluid density, ν is the flow velocity, and μ is the fluid viscosity. Dh represents the hydraylic diameter and, for rectangular channels, is given by

$$\mathbf{D}\_{\mathbf{h}} = \mathbf{2} \text{ a } \mathbf{b} / (\mathbf{a} + \mathbf{b}) \tag{2}$$

Where a and b are the cross sectional dimensions of the channel. The Reynolds numbers involved in the field of microfluidic device are usually below 100 (Gravesen et al., 1993), which prevent turbulent flow.

According the formula (1), the fluid density (ρ) and the fluid viscosity (μ) are constants. Re is depended on the flow velocity (ν) and the hydraylic diameter (Dh). For our microdevice with a 200µm wide and 100µm deep rectangle channel, the hydraylic diameter (Dh) calculated by the formula (2) is 133µm. Note that the blood density (ρ) is 1.05g/cm3, and the viscosity of blood (μ) is 3.115mPa.s. According to the formula (1), the Re in our experimental was 0.035~8.4, when the flow velocity (ν) is about 0.0008~0.024m/s. Therefore a typical laminar flow is found in our experimental. Lack of turbulent flow limits the effective mixing of the fluids under investigation. On the other hand, for the cell lysis on the microdevice based on the chemical method, it requires that each cell should be fully exposed to lysis reagents in microchannels, which is determined by mixing performance of the microdevice. In microdevice channels where flow is laminar in nature, the mixing is dictated by diffusion. Average time for a small portion of a fluid to diffuse a distance L can be estimated by

$$\mathbf{T\_D = L^2 / D} \tag{3}$$

Where D is the diffusion coefficient of the liquid. The above equation can be used to predict the order of time scale of mass diffusion. As it suggests, one can dramatically reduce the mixing time by reducing the diffusion length required for mixing or increasing the contact area between two different liquids while keeping the volume constant.

MEMS-Based Microdevice for Cell Lysis and DNA Extraction 27

Fig. 2. Numerical simulation results for the species concentration distributions at T-type mixing model (a), sandwich type mixing model with lined channel (b) and sandwich type mixing model with coiled channel (c) when νcell = 0.005m/s, νbuffer = 0.0025m/s. The color for "1.00e+00" and "0.00 e+00" in this figure refers to the concentration intensity of the pure cell sample and the pure lysis buffer, respectively. The software of FLUENT 6.2 was used to

The microdevice consisted of a silicon substrate and a glass cover. The fabrication process was shown in Fig. 1. A double-side polished n-type 0.01-0.1Ωcm silicon wafer of (100) crystal orientation was first deposited with 0.3μm thick of silicon nitride (Si3N4) as the mask for fabricating porous silicon by a chemical vapour deposition (CVD) method, and then with 0.3μm aluminum as the mask for deep reaction ion etching (DRIE) by a electron beam vapour deposition technology. The wafer was spin coated with a positive photoresist (AZ1500) and patterned. After the exposed photoresist was developed, the exposed aluminum was removed by phosphoric acid and then the exposed silicon nitride was removed by plasma etching. The wafer was etched in a deep reaction ion etcher (Adixen, AMS100) to produce rectangle channels of 100μm in depth. After the process of removing the remained aluminum, a porous silicon layer on the internal walls of channels was anodized in 10% HF electrolyte at 20mA/cm2 for 15min. After that, the silicon wafers were thoroughly rinsed in distilled water and oxidized at 1050°C for 1h. After oxidation process, the silicon nitride formerly patterned as the mask was removed by plasma etching for bonding. Finally glass covers (Corning Pyrex#7740), in which three holes had been drilled corresponding to the silicon substrate, were anodically bonded to the silicon wafers to form

the closed channel by a bonder (Suss, SB6), thus microfluidic devices are fabricated.

The microdevice was mounted onto the stage of a microscope with a CCD camera and a video monitoring system. Blood was introduced through the cell inlet by a peristaltic pump, while

**3.3 Online cell lysis and DNA extraction procedure** 

simulate.
