Introduction
Motivation for maximizing bonding strength of PDMS/PMMA for
Microfluidic devices are widely used in biology, pharmaceuticals, chemistry, food processing, and environmental engineering Basically, it is a science and technology of manipulating and controlling fluids, usually in the range of microliters (10 -6 ) to picoliters (10 -12 ) [1], in the networks of fabricated channels with cross- sections in the 1-500àm range, and small volume capacity
Figure 1-1 Characteristic length scales (approximate) of microfluidic and nanofluidic systems in relation to that of various biological entities [2]
Since 1979, microfluidic devices were first developed for gas chromatography [3] In addition to its advantages over conventional laboratory-scale experiments, microfluidics offers a variety of other benefits Figure 1-1 shows cells, molecules and other biological entities with characteristic length scales commensurate with their dimensions The inverse characteristic length scaling of the surface-area-to-volume ratio suggests that heat and mass transfer into or out of a chip can be enhanced as the device’s dimensions are reduced; other physicochemical interfacial phenomena not typically encountered at macroscopic dimensions can also be utilized Furthermore, separation may be conducted at smaller sizes more quickly and efficiently In addition, microfluidics offers integration capabilities such that the entire spectrum of benchtop laboratory protocols, from sample handling to reaction, separation, and detection, can be incorporated and automated onto a single chip in a manner analogous to that of chemical plant unit operations With advancements in large-scale mass microfabrication and nanofabrication, economies of scale can be utilized to produce inexpensive, disposable, and portable handheld devices that have the potential to revolutionize both personalized healthcare medicine through point- of-care diagnostics and high-throughput drug discovery through massive parallelization [2]
Initially, microfluidic device components were constructed of silicon substrates [3] Silicon is a relatively expensive, entirely opaque in the visible/UV region of the spectrum, and non-biocompatible During the fabrication of silicon devices, either subtractive (e.g., wet or dry etching) or additive (e.g., metal or chemical vapor deposition) techniques are used to form structures Moreover, silicon has a high elastic modulus (130–180 GPa) and is difficult to process into fluidic components such as valves and pumps b) Glass
After the initial emphasis on silicon, glass eventually became the preferred substrate Glass is compatible with biological materials, has low nonspecific adsorption, and is impermeable to gases; however, etching microchannels can be very difficult Working with glass materials is time consuming and requires highly sophisticated fabrication methods (e.g., anodic or fusion bonding) under tremendous temperatures and pressures
Figure 1-2 Silicon nanowire system for cardiac biomarker detection (a) Image of silicon nanowire (SiNW) device array chip, integrated with microfluidic system for fluid exchange Fluids are deposited into the acrylic well through the inflow tube on the left (red arrow) and removed from the outflow tube on the right (blue arrow) (b) Schematic showing the layout of the SiNW device array on the chip A total of 36 clusters of 5 nanowires each were available for use (c) SEM image of a cluster of nanowires Each nanowire is individually addressable by oxide- passivated metal contact lines running out to the edge of the chip [4] c) Polydimethylsiloxane (PDMS)
Since it was first developed by George M Whitesides in the early 2000s, polydimethylsiloxane (PDMS) has many excellent qualities as polymer, so they become the material widely used in the development of microfluidic devices [5, 6] PDMS provides high elasticity, with Young’s modulus (usually symbolized as E) varies wildly around ~1-100 N/m 2 (MPa) In practice the high elasticity of PDMS means that the surface feature can be stretched by amount much larger than their
5 dimensions PDMS also has high optical transparency, the transmittance of PDMS is 92~93% (water, by comparison, has 99.97%) The index of refraction of PDMS is 1.43 (at 20 0 C), very similar to glasses Low surface energy also makes it easy to replicate and peel-off from master lithography-molds at the nanometer scale Especially, PDMS is a versatile material for biological and cellular applications due to their chemical inertness, low polarity, and low electrical conductivity d) Thermal plastic
Thermoplastics are a desirable alternative to silicon and glass substrates for the fabrication of microdevices Compared to PDMS, thermoplastic possesses highly optically transparent and inexpensive, while providing good resistance to chemicals, high mechanical strength, and excellent biocompatibility [7] A variety of polymeric materials have been developed for the fabrication of microfluidic systems, including polycarbonate (PC), polystyrene (PS), cyclic olefin copolymers (COC), and polymethylmethacrylate (PMMA) The features of thermoplastic materials for use in microfluidic applications are summarized in Table 1-1 Physical, chemical, and optical properties of commonly used thermoplastics in microfluidics
Table 1-1 Summary of material properties for common and trending microfluidic thermoplastics such as: PMMA, COC, PS, PC [8]
Mechanical properties Rigid Rigid Rigid Rigid
A significant feature of thermoplastics is their ability to softened or fully melted and reshaped when heated above a characteristic glass transition temperature (Tg) While remaining chemically and dimensionally stable over a wide range of operational temperatures and pressures Together with a wide range of material properties, this feature makes thermoplastic highly adaptable for use as substrates for microfluidic applications
PMMA is the basic, main, and most representative member of acrylics This material is a commonly used for microfluidics since the early developments in
7 thermoplastic microfluidic system until now thanks to its wide availability in a variety of grades with high optical transmission in the visible wavelengths, with a refractive index of 1.50, the optical clarity and transparency of PMMA is very close to that of glass [9] PMMA is also less expensive, light, good solvent and chemical compatibility, and well-characterized molding parameters
1.1.3 Bonding PDMS/PMMA for fabrication microfluidic devices
The formation of microchannels and substrate bonding are the two key processes in the fabrication of microfluidic devices Researchers have used adhesive bonding, ultrasonic welding, and thermal fusion bonding in the homogeneous bonding of PMMA to PMMA While the homogeneous bonding between PDMS to PDMS can be achieved by creating a surface hydroxyl group via plasma treatment, corona discharges, or UV/ozone treatments, followed by thermal curing Achieving a heterogeneous and effective bonding between PDMS and PMMA substrates extended the applications of microfluidics For example, this heterogeneous bonding could create hybrid microvalves [10, 11], and micropumps [12] devices for cell culture and analysis [9] , multiplexed microbioreactors [13], micro lens arrays[14, 15], or extraction of bacterial DNA for diagnostics [16]
Direct bonding, such as modification-based bonding, has been commonly used to bond PDMS and PMMA substrates, and previously reported methods are shown in Table 1-1 Valchopoulou et al proposed a surface modification approach
8 to the bonding of PDMS to PMMA [17] Oxygen plasma is firstly used to activate the surface of PMMA with 5% 3-(Aminopropyl)triethoxysilane (APTES) The PDMS and activated PMMA surfaces are then exposed to oxygen plasma to create silanol groups on both surfaces Finally, the two surfaces are brought into contact to form Si-O-Si bonds The tensile strength of these PDMS-PMMA bonds was estimated at 1130 kPa Kim et al presented a PDMS-PMMA bonding process involving the modification of PMMA substrates using oxygen plasma in 5% APTES, followed by the corona discharge treatment of the PMMA and PDMS substrates [18]
In experiments, the tensile strength between PMMA and PDMS was 2500 kPa, with water burst pressure of approximately 350 kPa Note however that both of these methods require both oxygen plasma and corona discharge treatments as well as the high-temperature immersion of PMMA in APTES after treatment Sunkara et al adopted the similar approach, while decreasing the APTES concentration from 5% to 1% to create irreversible bonds at room temperature In experiments involving the bonding of PDMS to PMMA, tensile strength was 385 kPa and air bursting pressure was 528 kPa [19] Wu et al employed a silane coupling agent, 3- mercaptopropyltrimethoxysilane (MPTMS), to permanently bond PDMS to non- silicon-based materials, such as thermoplastics, copper, and aluminum The thermoplastics, metals, and alloys are treated using a 2% aqueous solution of MPTMS, while the substrates and PDMS are treated via corona discharge This approach to bonding PDMS to PMMA achieved tensile strength of 335.9 kPa [20]
Lee and Chung demonstrated an approach of combining two PDMS substrates using two silane reagents, such as APTES or (3-Glycidyloxypropyl)trimethoxysilane (GPTMS) [21] Basically, the two PDMS substrates are exposed to oxygen plasma for 1 min and then immersed in a 1% aqueous solution of APTES and GPTMS for
Microvalve and Micropump
Microvalves are the main components of several microfluidic systems Microvalves control the flow, timing, and separation of fluids within a microfluidic device, they are essential for systems with intricate functionality Micropumps, on the other hand, are utilized to reduce the amount of external hardware required to operate a microfluidic device by creating periodic and volumetric fluid movement on-chip
The bonding between PDMS and PMMA mentioned above have been used to exploit the advantages of PDMS membranes in the creation of micropumps and microvalves Quake et al introduced normally-open valves with a relatively simple multilayer PDMS construction [26] One layer contains channels for fluid flow, over which a cross-channel architecture is used to provide a “control channel.” Pressure applied on the control channel the membrane to deform, which subsequently interrupts the flow of liquid Note that the two layers were sealed via baking at 80 °C for 1.5 h, and the resulted working pressure of microvalve was less than 100 kPa Note that PDMS-PDMS microvalves are highly susceptible to distortion under high operation pressure
Figure 1-3 a) Micrograph and schematic of a PDMS pneumatic microvalve, (b)
3D diagram of an elastomeric peristaltic Quake’s pump [26]
Therefore, other researchers have realized the Quake’s microvalves in hard plastic chips by sandwiching a thin PDMS sheet between rigid substrates, such as PMMA [10, 24], polycarbonate (PC) [27, 28] Nonetheless, difficulties in bonding PDMS membranes to these materials has limited these microvalves to be applied in extended applications More recently, researchers have sought to create microvalves with different elastomers, such as polyurethane (TPU) [29, 30] and Viton® [31] as an alternative to PDMS Unfortunately, these materials require the use of highly sophisticated equipment with high-temperature bonding and achieved liquid bursting pressure of less than 345 kPa
The most fundamental issue of an integrated microfluidic system is transporting precisely a tiny amount of liquid to study reaction phenomena Moreover; the device infuses liquid into the microchips should following
15 characteristics: small size, low energy consumption, high precision, self-contained, high flexibility in terms of flow rate and adaptive different liquids or viscosities Micropump is capable of providing these inevitable requirements According working principle, micropumps can be categorized as non-mechanical without moving parts (i.e., non-membrane actuated) or mechanical actuator with moving parts (i.e., membrane actuated) [32-35] Among the extensively mechanical micropump, peristaltic micropumps have gained attention from researchers due to its low cost, high precision Numerous studies and developments have been conducted on peristaltic micro-pumps since the first version was investigated [36] In general, these micropumps typically feature diaphragms which is contained flexible membrane, a desired fluid flow is generated by peristaltic motion of deflection’s membrane Actuation principle of peristaltic have been demonstrated such as thermal [37, 38], motor [39, 40], piezoelectric[41-43], pneumatic [26, 44-46] Especially, pneumatic peristaltic micropumps (PPMs) are efficient ways to drive fluids through microchannels, simplicity in fabrication, and can be easily integrated with biosensors The pneumatically actuated micropumps were developed by Quake’s group has made a significant contribution to the field [26] PPMs are effortless to design as a module for portable and disposable microfluidic devices in applications such as automated immunoassay [47], PCR reaction [48], DNA attraction [10], screening tests and disease diagnostics [49]
The key factors being investigated on PPMs are materials, fabrication methods, and flow rates Importantly, flow rate of the pumps determine performance, type of liquids, the applications in microfluidics systems Currently, polydimethylsiloxane (PDMS) are the most widely applied in fabrication in PPMs due to its good biocompatibility, and simple fabrication process Unger et al introduced a normally-open micropump with a relatively simple multilayer PDMS construction [26] The PDMS layers were sealed via baking, the resulting working pressure of the micropump was less than 100 kPa; the flow rate of this micropump was 0.141 àL/min In accordance with this concept, PDMS sheets were bonded to fabricate PPMs by heating [50] or plasma treatment [51, 52] However; due to limitation of fabrication process, so that working pressure of these pumps less than
200 kPa, effecting flow rate less than 900 àL/min Note that PDMS/PDMS micropumps are highly susceptible to distortion under high operation pressure
Besides PDMS, polymethylmethacrylate (PMMA) is another common material for microfluidic devices because of its high rigid mechanical property, chemical resistance, easy fabrication and excellent transparent To fully benefit from PDMS and PMMA, Zhang et al proposed a hybrid micropump which fabricated by sandwiching a thin PDMS membrane and between a PMMA pneumatic channel and a PMMA fluid channel [10] The PDMS/PMMA bond was not strong, air force applied in microchamber was 60 kPa so that working flow rate was only 350.4 àL/min Another approach was introduced by Hsih et al., adhesive film was
17 implemented to bond irreversibly PDMS/PMMA The flow rate of 96 μl/min was achieved at working pressure of 138 kPa [12] Alternatively, one group of researchers replaced the PDMS membrane in the micropump with a polyurethane (TPU) membrane Unfortunately, the technique fabrication required highly sophisticated with high-temperature bonding, the flow rate of 382 àl/min was obtained as air actuation pressure of 100 kPa Microfluidic devices have recently been implemented using additive manufacturing; researchers have used resin material to fabricate micropump following Quake’s design [53] However, the printed-membrane lacked the flexibility and the devices were limited to low operation pressure of less than 100 kPa corresponding flowrate of 21.6àl/min These PPMs described above; although they can be made liquid move, the main drawback is low-pressure condition and limited flow rate
In PPMs, fluid is driven by multiple elastic membranes actuated by their respective pneumatic chambers Control of membrane movement is essential for pump performance as it significantly affects stroke volume The stroke volume is determined as the amount of working fluid swept through each chamber during a single actuation Meanwhile, the dead volume is defined as the volume that is not swept in the chamber Thus, to achieve a high-performance pumping rate, we should increase stroke volume while minimizing dead volume Easy to realize that stroke volume and dead volume then micro pump working defend on movement of membrane However, studies on the movement of the PDMS membrane under the
18 influence of air pressure are conducted in a stationary state (i.e., study under single pulsation of air force) [12, 50, 54], whereas there has been no direct experimental observation of movement membrane under high-frequency operating conditions Although many reports of PDMS micropump fabrication and applications are found, how to increase stroke volume, and behavior deformation of membrane under high- pulsed frequency are not clearly understood.
Blood plasma separation
Human blood is undoubtedly one of the most vital biological fluids utilized in medical laboratory diagnostics Blood includes approximately 45% (volume) of red blood cells (RBCs), white blood cells (WBCs), and platelets While blood plasma which takes up the remaining 55% of the total volume of the whole blood, consists of proteins, antibodies, fragments of DNA, and RNA [55] These analytes are essential for medical diagnostics therapeutic applications [56] Blood cells are typically eliminated prior to diagnostic testing because the biological components of blood can easily interfere with the operation of biosensors, regardless of sensor type, hence influencing the final assay results Although conventional approaches for blood processing as the whole human blood However, the centrifugation method induced by ultra-high-speed rotation is challenging to be well- trained, time-consuming, requires bulky equipment, and difficult to be integrated on microfluidic platforms [57, 58]
With the advent of microfluidics-based techniques, there has been an intense emphasis on efficient and powerful approaches for blood plasma separation Additionally, these devices are highly integrated and automated on a single chip, employing miniaturized
19 total analysis systems (TASs) or lab-on-a-chip (LoC) devices to minimize human intervention Numerous microfluidics-based approaches for blood plasma separation have been developed to satisfy the expectations of these applications In general, they can be divided into two groups: active separation and passive separation Active techniques rely on an external force field such as magnetic, acoustic or dielectrophoresis, whereas passive approaches rely solely on channel design and intrinsic hydrodynamic forces for functionality Passive separation techniques are typically preferred over active separation techniques because they typically do not require the integration of external force fields, whereas active separation techniques frequently require more complex devices and ancillary equipment, making them unsuitable for Point of care (POC) applications
Figure 1-4 Component of human whole blood
In recent years, inertial particle focusing has attracted great attention in subfield of microfluidics Inertial microfluidics capitalizes on hydrodynamic forces acting on cells to concentrate them within the flow These forces cause cells to migrate across streamlines and organize into equilibrium positions based on their size, microchannel structure, and flow rate [59-61] Besides the microparticle separation, the inertial platform would be suitable for blood plasma separation Several groups have proved that a spiral microfluidic device with a trapezoidal cross-section could efficiently isolate plasma from diluted blood [62-64] Blood is injected into a spiral microchannel with a constant flow rate, and, under the influence of inertial lift and Dean drag forces, blood cells are focused and concentrated near the inner wall around the vortex cores, while blood plasma is collected from the outer wall
To achieve the required flow velocities, the separating system typically uses an external syringe or a peristaltic pump to achieve steady blood flow rates inside the microchannel Although these external systems are effective, they are not ideal for àTAS or LoC systems that utilize relatively small volumes of blood In addition, the dead liquid volume in the connection tube is considerable, making it challenging to integrate functionality detection onto a single chip.
Objective and significance of this dissertation
As mentioned in Section 1.1.3, PDMS and PMMA are the most common materials for microfluidics The combination of those materials into one homogeneous or hybrid chip can help to exploit their advantages simultaneously However, in general, the irreversible bonding method described above requires
21 multiple processes, time-consuming, and substrate immersion at high temperature, and low bonding strength
Microvalve and micropump are key elements in microfluidic system Especially, applying elastic properties of PDMS and rigidly properties of PMMA to fabricate microvalve/micropumps following Quake’s design They more and more apply and develop in microfluidics However, limitation of bonded strength of between interface PDMS/PMMA made these devices worked under low pressure, effect to flowrate pass through devices, limited applications Many microfluidic applications could benefit from hybrid devices made of PDMS and PMMA; however, the existing bonding methods could not generate sufficient bonding strength and limit such applications
The effort of developing the proposed bonding PDMS/PMMA process is aims to address the following issues:
• Maximizing the interfacial bonding strength between PDMS and PMMA using highly reproducible and low-cost methods at room-temperature
• Based on this bonding technique, we propose a fabrication of microvalve/micropump process that combines the elasticity of PDMS and the rigid of PMMA that could be operated under extremely high air pressure and high flow rate operation
• Develop a micropump capable of transporting a precise and high throughput of a tiny of blood and separating blood plasma using the inertial focusing approach without external pumps or valves Moreover, the device should have the following characteristics: small size, low energy consumption, high precision, self-contained, high flexibility in flow rate.
Organization of the dissertation
As mentioned in Section 1.4, the main purpose of this thesis in present a novel bonding method for heterogeneous interfacial PDMS/PMMA, then apply it to the fabrication of microfluidic devices Thus, to present the study clearly, and logically, the subsequent chapters are organized as the following:
Chapter 2 is divided into three parts The first part is the literature review of methods used to directly combine PDMS/PMMA substrates The second part presents the current design, working principle, and the pneumatic microvalve/micropump limitations The final part is presented for currently designed chips for plasma separation using the initial microfluidic technique
Chapter 3 presents a systematical method, the Taguchi method, to maximize the bonding strength between PDMS-PMMA microfluidic devices by optimizing the process parameters at room temperature Various performed examinations to evaluate bonding quality, such as: burst tests with water and the air, tensile tests, high flow rate tests, and SEM tests, were also reported in this chapter
Chapter 4 presents the applying the proposed method to bond PDMS/PMMA for the fabrication of pneumatic microvalve/micropump Various examination methods were implemented to evaluate the performance of these microfluidic devices, such as measuring flow rates, back pressure, and working dead volume Then the micropump was integrated with a spiral microchannel with a trapezoidal cross-section area and used to extract plasma from human blood rapidly
Chapter 5 is the conclusion, limitations, and recommendation for this study
Literature review
Bonding PDMS/PMMA by surface modification
2.1.1 Surface chemical modification using one silane reagent
Initially, in 2009, Valchopoulou et al proposed a surface modification approach to the bonding of PDMS to PMMA [17]
Figure 2-1 Schematic of the process flow for bonding between a PMMA and a
Oxygen plasma is firstly used to activate the surface of PMMA with 5% 3-(Aminopropyl)triethoxysilane (APTES) The PDMS and activated PMMA surfaces are then exposed to oxygen plasma to create silanol groups on both surfaces The plasma treatment conditions were followed: pressure of 100 mTorr, a plasma power of 100 W, time for active PDMS: 48 seconds, and PMMA: 12 seconds Finally, the two surfaces are brought into contact to form Si-O-Si bonds The tensile strength of these PDMS-PMMA bonds was estimated at 1130 kPa
Kim et al presented a PDMS-PMMA bonding process involving the modification of PMMA substrates using oxygen plasma in 5% APTES [18], followed by the corona discharge treatment of the PMMA and PDMS substrates With try various combinations, the optimal conditions were figured out: PMMA surface treated with oxygen plasma by using the activated 3-APTES solution at
85 o C; after that, PMMA and PDMS active by corona discharged to bond In experiments, the tensile strength between PMMA and PDMS was 2500 kPa, with water burst pressure of approximately 350 kPa However, both of these methods require oxygen plasma and corona discharge treatments and the high- temperature immersion of PMMA in APTES after treatment
Figure 2-2 The surface modification process for bonding PMMA and PDMS
Wu et al employed a silane coupling agent, 3- mercaptopropyltrimethoxysilane (MPTMS) [20], to permanently bond PDMS to non-silicon-based materials, such as thermoplastics, copper, and aluminum The thermoplastics, metals, and alloys are treated using a 2% aqueous solution of MPTMS, while the substrates and PDMS are treated via corona discharge This approach to bonding PDMS to PMMA achieved tensile strength of 335.9 kPa
Figure 2-3 Schematics illustrating procedures for (a–f) surface modification with 2% MPTMS at room temperature on a non-silicon substrate, and (g–i) bonding of an oxidized non-silicon substrate with oxidized PDMS realized at room temperature for 10 min [20]
Sunkara et al further introduced a method: PMMA active by oxygen plasma for 1 min and placed in an aqueous solution of 1% v/v APTES for 20 min The PDMS also treated with an oxygen plasma (60 W) for 1 min And irreversible bonding of PDMS–thermoplastics was achieved at room temperature [19] The PDMS/PMMA device could be withstand large hydrodynamic flow with a per minute injection volume of nearly 24,000 times higher than the total internal volume of the microchannel
2.1.2 Surface chemical modification using two silane reagents
Tang et al employed the oxygen plasma treatment of thermoplastics and PDMS, followed by immersion in a 1% aqueous solution of APTES and GPTMS for 20 min When brought into contact, the PDMS and PMMA form an amine- epoxy bond [22]
Figure 2-4 (a) Surface hydroxylation of PDMS and plastic substrates by O 2 plasma treatment for 1 min (b) Aminosilane and epoxysilane anchoring on the
O 2 plasma-treated PDMS and plastic substrates, respectively (c) Conformal contact of the two substrates at room temperature for 1 h [22]
Zhang et al employed a chemical, Poly(dimethylsiloxane)-monoglycidyl ether terminated, in the fabrication of PDMS-PMMA microdevices [23] After
29 corona discharge treatment, PMMA substrates are treated using 5% aqueous solution of APTES at 80 °C, heated to 80 °C for 4 h with monoglycidyl ether terminated Finally, both the PDMS monolayer-coated thermoplastic and pure PDMS are plasma oxidized and brought into contact under pressure of 0.1 Mpa to produce an irreversible bond The bonding strength of PDMS to PMMA, under air pressure burst tests, is 586 kPa
Figure 2-5 Schematic illustrating bonding mechanism (a) Coating of low- molecular-weight PDMS on thermoplastic surface Bonding of (b) PDMS with thermoplastic or (c) two thermoplastics A thin layer of low-molecular-weight PDMS was formed on the surface of an aminosilane-functionalized thermoplastic via an amine–epoxy bond, and two substrates were finally bonded by forming a siloxane (Si–O–Si) bond [23].
Microvalve
Zhang et al introduced a pneumatic valve applying PMMA/PDMS/PMMA, the fabrication process sealing player based on UV
30 ozone cleaner method [10] By regulating the pressure in a deformation chamber on the pneumatic layer using a computer-regulated solenoid, the closing or opening of a valve can be controlled However, low bonding strength between players, the valve provides only 15.4 àL s −1 at 60 kPa fluid pressure
Figure 2-6 Cross-sectional views of a three-layer monolithic PMMA/PDMS membrane valve (A) and exploded and assembled illustrations of a single PMMA/PDMS membrane valve (B) a: PMMA pneumatic wafer; b: displacement chamber; c: PDMS membrane; d: PMMA fluidic wafer; e: pneumatic channel; f: fluidic channel [10]
Suzuki et at proposed a pressure-actuated membrane valves [24]; they presented the process for bonding PDMS and PMMA by utilizing the sol–gel method The working pressure was -50 kPa
Figure 2-7 (a) Schematic illustrations of the micro membrane valve composed of PMMA–PDMS–PMMA substrates and its actuation by applying positive or negative pressure (b) Micrographs showing the valve opening/closing with the applied pressures as indicated (c) Relation between the valve-controlling pressure (Pvalve) and the critical pressure of Pfluid to open/close the valve
Dashed line indicates Pvalve = Pfluid [24]
More recently, researchers have sought to create microvalves with different elastomers Pourmand et at introduced method for the fabrication of whole-thermoplastic chips with embedded pneumatic micro-actuators [29] The fabrication process consisted of laser micromachining and thermal fusion bonding for PMMA and thermoplastic polyurethane (TPU) The valve could block liquid flows with liquid pressures only of 30 psi
Figure 2-8 Design and characterization of the normally closed microvalve Schematics of the operation of the valve, (a) open, (b) closed, and (c) exploded view of the valve architecture (d) Image showing a fabricated chip under a microscope (e) and (f) Characterization of the valve operation versus different liquid pressures and actuation pressures [29].
Micropump
As mentioned in section 1.2.1, Unger et al introduced a normally-open micropump [26], as known the first version of pneumatic micropumps, with a
33 relatively simple multilayer PDMS construction The PDMS layers were sealed via baking at 80°C over 1 h, and the resulting working pressure of the micropump was less than 100 kPa; the flow rate of this micropump was 0.141 àL/min
Figure 2-9 A 3D scale diagram of an elastomeric peristaltic pump The channels are 100 μm wide and 10 μm high Peristalsis was typically actuated by the pattern 101, 100, 110, 010, 011, 001, where 0 and 1 indicate “valve open” and “valve closed,” respectively This pattern is named the “120°” pattern, referring to the phase angle of actuation between the three valves Other patterns are possible, including 90° and 60° patterns The differences in pumping rate at a given frequency of pattern cycling were minimal (B) Pumping rate of a peristaltic micropump versus various driving frequencies Dimension of microvalves = 100 μm by 100 μm by 10 μm; applied air pressure
Wang and Lee demonstrated a micropump with a serpentine-shaped (S- shaped) microchannel that enables single-source actuation of three, five, or seven
34 actuation chambers [51] Compressed air deflects the membrane with a phase lag as it travels through the microchannels, the process which enables peristalsis The chip size was restricted; however, the serpentine-shaped air channel fills almost instantly, greatly reducing pumping rate The pneumatic micropump was made of PDMS by using plasma bonding technique A maximum flow rate of 108 àL/min was obtained at a frequency of 10 Hz and an air pressure of 172.4 kPa
Figure 2-10 A schematic illustration of the S-shape pneumatic microchannel with the integrated microvalve and microflow sensor Note that the pneumatic microvalve controls the injection of the fluid flow and the microflow sensor measures the pumping rate The phased deflection of the membranes located at the intersections of the S-shape microchannel and the fluidic microchannel generates a peristaltic effect which drives the fluid along the microchannel Parameters X and Y denote the length and separation, respectively, of the coils of the S-shape microchannel [51]
Furthermore, Yang et al reported the pneumatic micropump integrated with a normally closed valve that was capable of creating a high pumping rate [52] The normally closed valve was a PDMS-based floating block structure located inside the microchannel, which was activated by air pressure and created by the peristaltic motion of the PDMS membranes The results showed that the pumping rate was as high as 900 àL/min These micropumps mentioned above
35 were fabricated by using PDMS material; however, PDMS-PDMS micropumps are highly susceptible to distortion under high operating pressure
Figure 2-11 a Simplified fabrication process of the micropump based on SU-
8 lithography and PDMS replication The SEM images of the b SU-8 template and c PDMS inverse structure after replication [52]
Alternatively, researchers have sought to create micropumps with different elastomers, such as polyurethane (TPU), as an alternative to PDMS [30] However, these materials require highly advanced equipment associated with
36 high-temperature bonding The maximum flow rate of 382.497 àl/min was obtained at air actuation pressures of 160 kPa
Figure 2-12 Fabrication, performance and characterization of the micropumps (a) The image of a fabricated micropump; (b) micropump actuation pattern [30].
Separation of blood plasma using size-based particle separation with
Nivedita et al introduced the device to achieve blood cell separation, exhibit high separation efficiency (∼95%), with diluted blood [65] In Figure 2-13, the microchannels were fabricated in PDMS using the conventional soft lithography process The design chip included a three outlets system with a cross- section of 250 μm × 75 μm The optimal condition for blood plasma separation at 500× diluted blood, and flow rate of 1.8 mL/min
Rafeie et al demonstrated a spiral microchannel with the trapezoidal cross- section by which can enhance the cells focusing abilities in the microchannel The diluted blood samples achieve the 100% purity of separated plasma under a 1.5 mL/min flow rate [63] All the microchannels made of PDMS using a soft lithography technique from an aluminum mold The final device was irreversibly
37 bonded with a thick PDMS slab using an oxygen plasma method Figure 2-14 shows schematic illustration of the blood plasma separation using this design
Figure 2-13 (a) Image of design (b) Bright field image of the outlet system (c)
At 1.8 ml/min flow rate (F1), two focused streams are observed, the narrow stream in the middle of the channel is formed by 7.32 μm particles and the broad stream near the inner channel wall is the composite of three streams of
10, 15, and 20 μm particles (d) Fluorescent image of the focused streams of all three particles, 10 μm, 15 μm, and 20 μm in diameter, at the flow rate of 2.2 ml/min (F2) (e) Normalized focusing position of particles (x is the distance of the focused stream from the inner channel wall, and w is the width of the channel) as function of De [65]
Figure 2-14 Schematic illustration of the blood plasma separation using a spiral channel with a trapezoidal cross-section Under the influence of inertial lift and Dean drag forces, blood cells are focused and concentrated near the inner wall around the vortex cores, and blood plasma is collected from the cell-free region (outer wall) The bifurcation point is positioned closer to the inner wall to facilitate a large volume of plasma collection [63]
Robinson et al informed a microfluidic device contain the main spiral microchannel followed by two secondary spiral microchannels [66] The function of the two secondary microchannels is to filter out the blood cells with a similar flow velocity to the main microchannel, as shown in Figure 2-15 They finished the experiments with 100x diluted blood and received 99% separation efficiency with the device with secondary spiral microchannels and 55% efficiency with the single main spiral microchannel with the optimal configured flowrate of 1.25 mL/min The rectangular dimensions of 500 àm for width and 60 àm were designed, while the dimensions of the secondary spirals, reduced to 250 μm by
60 μm The chips were made by using PDMS bonded with glass slides
Figure 2-15 Brightfield images of the channel while filtering blood (a)
Solidworks model of the spiral channel indicating areas of imaging (b) Image of the first bifurcation with blood cells not completely filtered; the shadow boxed in red shows the cells passing the bifurcation to be filtered again by the second spiral (c) The second bifurcation where the cells passed by the first spiral are removed [66].
Maximizing bonding strength of PDMS/PMMA
Materials
The commercial transparent PMMA sheets with a thickness of 2 mm were purchased from Cho Chen Ind Co., Ltd (Tainan, Taiwan) PDMS and a curing agent were purchased from Sylgrad 184 (Dow Corning, USA) 3-
Aminopropyltriethoxysilane (APTES, 99%) and 3- glycidoxypropyltrimethoxysilane (GPTMS, 98%) were purchased from Sigma-Aldrich (St.Louis MO, USA) Isopropyl alcohol was purchased from Echo Chemical (Taiwan).
Bonding procedure
We revised the method proposed by Tang et al [22] for bonding between PDMS and PMMA substrates and carefully characterized the performance in this study, with the purpose to maximize the interfacial bonding strength between PDMS and PMMA using highly reproducible and low-cost methods at room- temperature Figure 3-1 presents a schematic diagram showing the process used to create irreversible bonds between PDMS and PMMA The layout was designed
40 using SOLIDWORKS and post-processed for subsequent machining using the Siemens NX A high-speed computer numerical control (CNC) micro milling machine (HM4030L, CHMER, Taichung, Taiwan) was used for the direct engraving of S-shaped microchannels (length of 70 mm, width of 500 àm, depth of 500 àm) on PMMA substrates using a two-flute end mill with a diameter of
500 àm (NS TOOL Co., Ltd., Japan) Two through-holes (1.2 mm in diameter) were fabricated in the upper layer for connections
Figure 3-1 Schematic illustration showing the process used to bond PDMS to
We adopted the following parameters to optimize the surface roughness and restrict the burrs that remained in the microchannels after engraving: spindle speed of 20,000 rpm, feed rate of 300 mm/min, and depth of cut of 10 àm [67] After manufacturing, the substrates were thoroughly washed using 10% isopropyl alcohol (Echo Chemical Co., Ltd., Taiwan) under sonication for 5 min and then dried On the other side, the PDMS substrates were prepared by combining an elastomer and a curing agent (Sylgrad 184, Dow Corning, USA) at a volume ratio of 10:1 The mixture was degassed for 15 min before being poured into custom
Micro milling Oxygen plasma Oxygen plasma
PDMS PDMS-PMMA chip PMMA
41 designed molds with dimensions precisely matching the machined PMMA substrates (height×width×thickness; 54 mm × 42 mm × 2 mm) The substrates then underwent thermal curing at 80 °C for 4 hours The two substrates were activated simultaneously using pure oxygen plasma The original polymer chains on the PMMA and PDMS surfaces were broken using a plasma machine (PDC- 32G, Harrick Plasma, Ithaca, NY, USA) to form hydroxyl groups Immediately after plasma activation, the modified PDMS substrate was soaked for 20 min in an aqueous solution of silane reagent (APTES, 97%, Sigma-Aldrich) to generate amine functionalities, whereas the modified PMMA was soaked for 20 min in an aqueous solution of epoxysilane (GPTMS, 98%, Sigma-Aldrich) to create epoxy functionalities After thorough drying, the two substrates were brought together using a 1 kg weight at room temperature to form a tight amine–epoxy bond Finally, the chip was attached with two PMMA machined connectors using glue.
Experimental setup
We conducted burst pressure tests to characterize the maximum bonding strength in microfluidic PMMA channels bonded to PDMS Figure 3-2(a) shows the schematic of the testing system, while Figure 3-2(b) shows the real testing system, the bonding strength was characterized using a fluidic system including a syringe pump for the injection of red food dye, a pressure meter (PS-9303 SD, Lutron Electronic Enterprise Co., Ltd., Taiwan), and a pressure sensor (PS100–
20 bar, Yalab, Taiwan) attached to a computer for data collection We also designed an aluminum chip holder to enable the use of a camera to identify
42 instances of leakage A three-way connector was used to ensure the transfer of pressure into the microchannel, while allowing detailed measurements of cumulative pressure within the microchannel With the outlet blocked, liquid was injected into the microchannel at progressively higher pressures until failure The bonding strength is described as the pressure at which interfacial separation was observed This testing method is well suited to applications requiring chip longevity and stability
Figure 3-2 The system used to measure the burst pressure of the bonded microfluidic chips
Note that each stage of the testing process was repeated three times After understanding the maximum bonding strength, leakage tests were conducted on the hybrid microfluidic devices, and particularly those that require a high flow rate associated with high operation pressure
Design of experiment
The proposed bonding strategy involves a number of parameters requiring optimization: GPTMS concentration (A), APTES concentration (B), radio frequency power (C), plasma treatment duration (D), and barrel reactor pressure (E) As shown in Table 3-1, we examined three power levels (6.8, 10.5, and 18 W) and three pressure levels (400, 600, and 800 mTorr) under exposure periods ranging from 60 s to 120 s In forming an amine-epoxy bond, we varied the APTES and GPTMS concentrations as follows: 1%, 3%, and 5% Normally, a total of 243 experiments would be required to thoroughly evaluate the effects of all parameter combinations on bonding performance In the current study, we employed the Taguchi experimental design process to identify the primary parameters affecting PDMS/PMMA bonding efficiency using a minimum number of experiments In Table 3-2, we employed a standard orthogonal array in the design of 18 parameter combinations
Table 3-1 Five primary variables and their corresponding values influencing bonding strength
Note that each three replicate experiments were conducted for each parameter combination Finally, the parameters were optimized with the goal of maximizing interfacial bonding strength for PDMS-PMMA microfluidic devices Table 3-2 presents results, the standard deviation (s) and signal-to-noise ratios (S/N) obtained at three measurement points Higher bonding strength is desirable; therefore, we calculated the signal-to-noise ratio (S/N) using Eq (1), where n = 3 indicates the number of times each combination was assessed, and yi 2 indicates the summed bonding strength of that combination
Table 3-2 Results of eighteen experiments based on factors and levels listed in, s is standard deviation and S/N is signal to noise ratio
Figure 3-3 shows types of broken chips corresponding to each condition For example, as shown in Figure 3-3(a), the condition No.6 has the lowest bonding strength; at the pressure of 55 kPa, liquid delaminate the entire interfacial PDMS/PMMA In Figure 3-3(c), at the pressure at 101 kPa, the liquid delaminates only some points along with the microchannel (the condition No.15) Figure 3-3(b) and Figure 3-3(e~h), the chip was observed with no leakage, even though the pressure in the device was over 600 kPa (the condition No.11)
Figure 3-3 Results of burst tests on devices fabricated using the various parameter combinations; (a) No.6 shows failure at the PDMS–PMMA interface under pressure of 55 kPa; (b) No.11 in which chip remained intact under pressure of 622 kPa; (c) No.15 showing local failure under pressure of
101 kPa; (d) enlarged photo of No 15; (e)-(h) Microscopic images of areas
Factor analysis
Factor analysis was used to identify the parameters with the most and least profound effects on the strength of bonds between PDMS and PMMA
Table 3-3 Factor analysis based on average S/N ratios
Figure 3-4(a~e) plots the average S/N ratios listed in Table 3-3 and the corresponding average bonding strengths The bigger-is-better objective function was used to analyze the results, due to the fact that the objective here was to maximize bonding strength Thus, a higher average bonding strength is indicated by a higher average S/N ratio
Factor E presented the largest range (6.530), indicating that the pressure (or flow rate) of pure oxygen inside the chamber had the most pronounced influence on the PDMS/PMMA bonds Factor B presented the smallest range (2.911), indicating that it had the least influence on bonding strength
Figure 3-4 S/N ratios and bonding strengths corresponding to five main factors and three levels: (a) GPTMS concentration; (b) APTES concentration, (c) radio frequency power, (d) plasma treatment duration, (e) pressure within barrel reactor
Thus, Figure 3-4(a~e) can be used to determine the optimal bonding parameters in terms of maximizing PDMS/PMMA bonding strength, meaning the highest bonding strength of 622 kPa can be realized with parameters of A1 (GPTMS concentration of 1%), B2 (APTES concentration of 3%), C1 (RF power of 6.8 W), D1 (exposure time of 60 s), and E3 (oxygen pressure of 800 mTorr), which corresponds to No 11 in the Table 3-2 With the standard deviation of 6.53 and coefficient of variation (COE) of 0.01, the resulted parameters are reliable and robust
The concept underlying the proposed approach to bonding PMMA to PDMS is to incorporate Si-containing functionalities onto the PMMA surface As shown in Table 3-3, pressure within the chamber had the most profound overall effect The bonding strength of PDMS and PMMA is maximized when the pressure of oxygen is 800 mTorr in the plasma generator chamber (No 11 in Table 3-2), meaning pumping more oxygen into the plasma chamber is beneficial to the bonding strength On the contrary, the bonding strength decreases gradually if the oxygen pressure inside the plasma chamber is reduced, as shown in Figure 3-4(e) The results are consistent with the results of Bhattacharya et al [68] Figure 3-4(c) and Figure 3-4(d) present plots of S/N ratio and average bonding strength as a function of RF power and exposure duration These two factors present similar trends, higher power or longer exposure duration would lower the bonding strength After plasma treatment, APTES is used to bind activated PDMS to form amine functionalities, whereas GPTMS is used to bind activated PMMA
50 to form epoxy functionalities Figure 3-4(a) and Figure 3-4(b) show an interesting phenomenon, which seems that two chemicals are critical to the bonding between PDMS and PMMA substrates, but they have opposite influence to the bonding strength The lowest bonding strength was obtained under a concentration of 3% GPTMS, while the highest bonding strength was obtained under a concentration of 3% APTES Nguyen at al [55] reported that an APTES concentration of 5% with reaction duration of 20 min produced the highest density of surface amines Nonetheless, the highest durability in the current study was achieved when APTES was combined with GPTMS at a concentration of 3% and 1% respectively.
Conformation test
To confirm the bonding strength of PDMS/PMMA, we take the bonded chips with burst test with the air pressure, tensile test and high flow rate test, and high-density microchannels
3.6.1 Burst test with the air
The chip was tested with the air The schematic is shown in Figure 3-5 First, the outlet block, open the valve when measuring the pressure in the microchannel by the pressure meter And the pressure can withstand up to 770 kPa (maximum of the pressure supplied by the air compressor) Then, we removed the air tube and injected dye food water into the microchannel The chip can operate without leakage The magnitude of the air pressure in the microchannel is more significant than in all previously published studies
Figure 3-5 The system is used to measure the burst pressure with the air of the bonded microfluidic chips
Figure 3-6 (a) PDMS/PMMA chip with are bonding by 22 mm × 22 mm for the tensile test; (b) The microchip was set on a tensile test machine; (c) The tensile test of the PDMS/PMMA bonding at a force 1.5 kN
With these optimized parameters for the high bonding strength, the bonded hybrid microfluidics were tested with standard tensile tests (INSTRON 3365, USA), as shown in Figure 3-6(a)~(b) The experiment results showed that the tensile strength of the PDMS/PMMA assemblies formed by the optimal bonding process exceeding 3,000 kPa, in Figure 3-6(c)
Many applications require compact microfluidic devices, which requires the ability to manufacture microfluidic devices with a minimum space between adjacent microchannels, which results in high back pressure inside the microchannel and delaminate the bonded chips due to the insufficient bonding strength
Figure 3-7 (a) Layout of microfluidic device used to determine the effectiveness of high-density microdevices The chip included main microchannel (No 1) (width of 200 àm and depth of 100 àm), four secondary microchannels (width and depth of 100 àm) with various spacing between adjacent microchannels (No 5) 200 àm, (No 4) 100 àm, (No 3) 50 àm, and (No 2) 30 àm; (b) the chip after injecting red dye ink and compared to a ten dollar coin; (c) Microscopic image of sub-microchannels (No 2), (No 3), (No 4), (No 5) shown in (b); (d) Microscopic images of marked areas (I), (II),
We manufactured microfluidics chips specifically to determine the highest microchannel density that could be achieved using the proposed PDMS-PMMA assembly method As shown in Figure 3-7(a), the test platform included a main
54 microchannel with a width of 200 àm and depth of 100 àm (No 1) as well as four secondary (i.e., narrower) microchannels with a width of 100 àm and a depth of
100 àm The distance between each parallel microchannel was varied as follows:
200 àm (No 5), 100 àm (No 4), 50 àm (No 3), or 30 àm (No 2) Testing involved injecting ink solution into the main microchannel (No 1) using a peristaltic pump (DGS, YZ1515X, Shenchen, China) at a flow rate of 1 mL min −1 , as shown in Figure 3-7(b), and four secondary microchannels (No 2 to No.5) were investigated
The fabricated high-density compact microfluidic chips with rectangular parallel microchannels, and the results are shown in Figure 3-7(c) and Figure 3-7(d), where Figure 3-7(c) shows the four secondary microchannels shown in Figure 3-7(b) and Figure 3-7(d) shows the enlarged areas shown in Figure 3-7 (c) with marks of I, II, III, and IV Note that spacing between microchannels ranged from 30 àm (I in Figure 3-7 (d)) to 200 àm (IV in Figure 3-7 (d)) For channels of a given length, reducing the distance between channels to 30 àm would significantly reduce the chip area by nearly 50% For example, a passive micromixer with an area of 2 cm ì 2 cm with a channel width of 100 àm and gaps of 30 àm would provide a total length of more than 3.0 m for uniform and sufficient mixing Water containing red food dye was pumped through the microchannels at a flow rate of 1 mL min −1 and monitored under a microscope to identify any leakage between microchannels in all cases, note that the flowing velocity was 416 mm s −1 in the secondary microchannel (cross-section of
100 àm ì 100 àm) No leakage was observed even when the gaps were reduced to 30 àm Thus, it appears that this bonding technique could be used to create long and narrow microchannels within a tiny area, which significantly reduces the size of microfluidic device
Scanning electron microscopy (SEM) was also used to investigate the cross-sections of the PDMS-PMMA after experiments and the results are shown in the Figure 3-8(a) and Figure 3-8(b) It is clear that an excellent bonding between the two heterogeneous materials still remains, even when the distance between adjacent channels was very narrow (30 àm for Figure 3-8(a) or 50 àm for Figure 3-8 (b)), was realized
Figure 3-8 SEM cross-sections showing excellent bonding between PDMS- PMMA with gaps of (a) 30 àm, and (b) 50 àm after experiments
To provide an extremely high flow rate for microfluidic devices, a peristaltic pump was used In this test, four test microfluidic chips with long and identical length of 70 mm were fabricated with different cross-sections for
56 experiments to understand: (1) any leakage happening under extremely high flow rates; (2) recording the accumulated pressure inside the microchannels
Figure 3-9 (a) System used to conduct experiments of high flow rates, including a peristaltic pump, two reservoirs, a connector, a pressure sensor, a pressure meter, and the microchip; (b) Experiment setup
The experiment for test flow rate as shown in the Figure 3-9 The width and depth of the crossed channels were as follows: 500 àm ì 500 àm,
300 àm ì 500 àm, 200 àm ì 500 àm and 100 àm ì100 àm The flow rate was controlled by the peristaltic pump from 5 mL min −1 to 190 mL min −1 while a pressure sensor was linked to record the pressure inside the microchannel in terms of flow rates
Figure 3-10(a-d) shows the red dye solution injected into the microchannels by the peristaltic pump at flow rates from 5 mL min −1 to
190 mL min −1 without any delamination or liquid leakage, and Figure 3-11 shows the recorded pressure inside the microchannel when different flow rates were used in the different size of microchannels Several conclusions can be drawn from the Figure 3-11: (1) the microchip could withstand pressure from 8 kPa (i.e
5 mL min−1, cross-section of 500 àm ì 500 àm) to 442 kPa (i.e 190 mL min−1,
57 cross-section of 100 àm ì 100 àm) without any leakage or delamination; (2) under the same flow rate, the pressure inside the smaller microchannel was significantly higher than the larger microchannel For example, 375 kPa was needed to generate a 50 mL min −1 flow rate in a 100 àm ì 100 àm microchannel while only 100 kPa was needed for 500 àm ì 500 àm microchannel; (3) when the pressure was smaller than 300 kPa, the pressure is proportional to the flow rate and linear, which aligned with the Poiseuille equation in which the flow rate is proportional to the pressure drop But all the curves among all size of microchannels went nonlinear when the pressure was roughly above 300 kPa, which was due to the limiting pressure drop of 300 kPa this peristaltic pump could provide, meaning the actual flow rate inside the microchannel was not the flow rate showing on the panel of the peristaltic pump In other words, those curves below the pressure of 300 kPa is reasonable and useful to understand the real flow rate inside the microchannels of various cross-sections Note that 300 kPa is still far lower than the bonding strength achieved in this study (622 kPa, the red solid line on the top of Figure 3-11), therefore even much higher flow rate can still be used in those microfluidics devices which followed the bonding protocol reported herein; (4) the bonding strength achieved herein allows a microchannel with a cross-section of 500 àm ì 500 àm to withstand an extremely high flow rate of
Design and fabrication of microvalve
The proposed valve actuation mechanism is based on the pressurization/depressurization of an air chamber containing a movable elastomeric membrane The design of the normally open valve is shown in Figure 4-1 The microvalve comprised a liquid chamber with a hemispherical shape embedded within a PMMA layer as well as a PDMS membrane and a PMMA cover as a control chamber for membrane actuation
Figure 4-2 Exploded view of valve architecture with schematic showing the operation of the valve in (a) open stage and (b) closed stage
The thickness of the PDMS membrane was 200 àm prepared by spin coating All layers were bonded together using the bonding method mentioned above Figure 4-2 shows the cross-section view of closed microchannel when the PDMS membrane was pressurized and sealed the microchannel
Figure 4-3 Schematic illustrations of the fabrication process
Liquid chamber Valve open Valve closed
The PMMA/PDMS/PMMA microvalve fabrication process is depicted in Figure 4-3 Firstly, the microchannels were fabricated by micromachining in Figure 4-3(a) The layouts were designed and post-processed for subsequent machining using SOLIDWORKS and Siemens NX software A high-speed computer numerical control (CNC) micro milling machine (HM4030L, CHMER, Taichung, Taiwan) was employed to engrave the pneumatic chamber and the microchannel on PMMA substrates Micro-end-mill and micro-ball-mill tools were used in experiments (NS TOOL Co., Ltd., Japan) We adopted the following parameters to optimize the surface roughness and restrict the burrs that remained in the microchannels after engraving: spindle speed of 20,000 rpm, feed rate of
300 mm/min, and the depth of cut of 10 àm The flexible PDMS membranes is employed in the architecture of the micropumps The PDMS membrane was prepared by combining the elastomer and the curing agent at a volume ratio of 10:1 The mixture was degassed for 15 min The thickness of the PDMS membrane was 200 àm which prepared by spin coating The membranes then underwent thermal curing at 80 °C for 2 hours After manufacturing, both were thoroughly washed using 10% under sonication for 5 min and then dried The bottom PMMA and the PDMS membrane were directly bonded using chemistry bonding that follow the proposed method
Figure 4-3(b) shows PMMA and PDMS surfaces were broken using a plasma machine to form hydroxyl groups Immediately after plasma activation, the modified PDMS substrate was soaked in an aqueous solution of 3% APTES,
63 whereas the modified PMMA was soaked in an aqueous solution of 1% GPTMS, as shown in Figure 4-3(c) After thorough drying, the two substrates were brought together using a 1 kg weight at room temperature to form a tight amine–epoxy bond in 1 hour At this point, fluidic access holes were punched on PDMS surface using a Uni-Core™ Puncher Next, in the previous PDMS-PMMA assembly was connected with the cover PMMA to make fully PMMA/PDMS/PMMA microvalve using the bonding method mentioned above (Figure 4-3(d)~f)) Finally, the device was attached with the PMMA machined connectors using glue for world-to-chip interconnection.
Characterization of microvalve
Figure 4-4(a) and Figure 4-4(b) illustrate the experiment setup used to characterize the microvalve Various gas actuation pressure levels were applied to the actuation chamber, while liquid was injected through the microvalves at various flowrates by the peristaltic pump Valve actuation was implemented using a custom-made controller system comprising two hand-operated pneumatic valves A precise gas regulator was used to adjust the actuation gas pressure To estimate the level of microvalve openness under various pressures, the volume of the liquid passing through the microvalve was determined by measuring the length of the liquid column for 1 min (Figure 4-4(c))
Figure 4-4(a) Schematic illustration of microvalve performance test; (b) Image of experiment setup; (c) Enlarged image of microvalve, the diameter of the tube is 0.38 mm
The opening and closing of the microvalve is based on deformation of a PDMS membrane Thus, the microvalve must be able to simultaneously withstand the pressure created by the liquid flow as well as to be rapidly actuated by air compression The operation of the microvalve is presented in https://youtu.be/itZRxCL8DZ4, and it clearly shows that this microvalve has rapid response and capability to completely close the microchannel when the liquid pressure is as high as 402 kPa Figure 4-5(a) shows the microvalve is open (corresponding to Figure 4-2(a)) and Figure 4-5 (b) shows the microvalve is
Micro valve Pressure actuator Ruler
65 closed (corresponding to Figure 4-2(b)) Figure 4-6 shows the experiment results from a microchannel with cross-section of 200 àm ì 60 àm, in which the liquid pressure means the pressure inside the liquid microchannel and the actuation air pressure means the pressure inside the air microchannel The actuation air pressure is controlled and maintained from 100 kPa to 700 kPa with an increasement of 50 kPa for all experiments
Figure 4-5 The microscope's photo shows the valve operation in (a) open stage and (b) closed stage
Figure 4-6 Openness of microvalve versus different liquid pressure ranging from 228 kPa to 402 kPa and actuation air pressure ranging from 100 kPa to
Openness percentage in the Figure 4-6 refers to the ratio of the measured volume of the liquid passing through the microvalve under various actuation air pressure to the measured volume of the liquid when the valve is fully opened over a period of 1 min Figure 4-6 clearly showed that: (1) the bonding strength reported herein can be successfully applied and made the microvalve working, no liquid passed through the microvalve even at a high liquid pressure of 402 kPa; (2) a higher actuation air pressure is needed to completely close the microchannel when a high liquid pressure is applied; (3) it is interesting to see that the openness percentage of microvalve is linearly proportional to the actuation air pressure, meaning only controlling the actuation air pressure can potentially be used to
67 control the flow rate of liquid; (4) by controlling the openness of the microvalve, we can efficiently control the flow rate in the microchannels; (5) the microvalve is suitable for applications which require high flow rate and high-pressure operation To the best of our knowledge, the air pressure and liquid pressure in our microvalves exceeded those of all previous devices made of thermoplastic, PDMS, or resin materials
We subsequently extended this concept to an integrated flow control system in a compact configuration for microfluidic devices, including eight liquid microchannels and seven air channels shown in Figure 4-7(a) The liquid channels could be opened/closed in any order and any combination with rapid response to enable the transport and sequential merging of various liquids, as shown in Figure 4-7 (b)~(e) Initially all the microvalves were closed, and the red dye solution was injected from the bottom microchannel Then the microvalve was opened sequentially and counterclockwisely, then the red dye solution flowed through each microchannel counterclockwisely in sequence This idea is scalable and applicable to high throughput applications It is particularly well suited to microfluidic-based assay systems requiring high accuracy and the easy handling of liquids
Figure 4-7 Microvalve system including eight liquid microchannels and seven air channels: (a) the area of this compact microvalve system was within diameter of 4 mm; (b) channel No 2 open; (c) channel No 2 and No 3 open; (d) channel No 2, No 3, No 4, No 5, No 6, and No 7 open; (e) All channels open.
Design and fabrication of micropump
The design of the normally open micropump is based on the concept of reciprocating series diaphragm microchambers, fluids are typically driven by multiple elastic membranes actuated by their corresponding pneumatic chambers
Figure 4-8(a) shows the structure of the PPM consists of three layers: a rectangular cross-section microchannel subsequence hemispherical shape chambers embedded within a PMMA layer (bottom player), an elastomer PDMS membrane, and a PMMA cover as a control chamber for membrane actuation Figure 4-8(b) illustrates the structure cross-section of the micropump Figure 4-9 (a)~(d) illustrate the working principle of the proposed micropump
Figure 4-8 (a) Design of the micropump; (b) Schematic showing micropump chambers and pneumatic actuation chambers
In the chamber 1, the PDMS diaphragm is deflected when the actuator chamber is subjected to external pneumatic force If the applied pneumatic force is sufficient, the PDMS membrane is completely closed to the half-spherical chamber area preventing fluid flow, as shown in Figure 4-9(a) After a specific time, the pneumatic force will be released and the subsequent chamber will be activated immediately (Figure 4-9(b)) Similarly, other membrane diaphragms can be employed to control fluid behavior, as shown in Figure 4-9(c)~(d)
Figure 4-9 The working principle of the proposed micropump (a) The first chamber closed (b) The second chamber closed (c) The third chamber closed (d) The fourth chamber closed The green arrows denote valve opening and the red arrows denote valve closure, blue arrows denote overall direction of fluid flow
The PMMA/PDMS/PMMA micropump fabrication process is depicted Figure 4-10 The process is similar to fabricating microvalves, as presented in section 4.1.2
Figure 4-10 Schematic illustrations of the fabrication process; (a) Fabrication of microchannel on PMMA substrates by using CNC machine; (b) Oxygen plasma treatment on PMMA substrates and PDMS membrane; (c) PMMA surface modification with GPTMS; and PDMS surface modification with APTES; (d) Oxygen plasma treatment on PMMA substrate and PMMA/PDMS chip; (e) PMMA surface modification with GPTMS; and PMMA/PDMS surface modification with APTES; (f) Micropump
All the PMMA substrates were fabricated using the micro-milling method; the method provides excellent quality of precision, accuracy, and good surface roughness Furthermore, the samples were captured using the profilometer microscope to investigate the surface roughness and structure of the chamber Figure 4-11(a) presents a 3D photo and surface roughness of machined channels achievement as Ra=0.45 àm; the milled surfaces are acceptable for almost microfluidic applications Figure 4-11(b) shows the SEM of the actuator
72 diaphragm cross-sections after bonding There is an excellent interfacial bonding between the three heterogeneous players Figure 4-11 (c) illustrates the noticeable peristaltic pump with inlet, outlet, and connected air tube actuators
Figure 4-11 (a) Three-dimensional of the PMMA half-spherical chamber with the depth chamber of 300àm (b) SEM of PMMA/PDMS/PMMA interfacial of micro chamber (c) Visualization of the pneumatic peristaltic micropump.
Characterization of micropump
The deflection of the PDMS diaphragm in response to air pressure is a critical factor affecting the performance of the PPMs Two experiments were implemented to study behavior of the PDMS diaphragm First, the center deflection and the radius curvature were measured under various air pressures
73 when the top membrane was free, as shown in Figure 4-12(a~b) The test chips contained pneumatic chambers in the shape of circles with a diameter of 2000 àm, the same as the diaphragm diameter Investigations of the PDMS membrane were carried out by using a calibrated pressure source The center deflection and the radius of curvature of the deformed membrane were captured at different applied pressures by the 3D laser profilometer machine (Keygen, USA), which varied from 50 kPa to 200 kPa
Figure 4-12 (a) Schematic of experimental setup for the measurement deformation of PDMS membrane under applied pressures, the system consists of: an air compressor, a regulator to control pressure, a hybrid PDMS-PMMA chips containing an air channel and a PDMS membrane, an objective lens of the 3D laser profilometer machine (b) Experimental setup
Objective lens of 3D laser profilometer machine
PDMS-PMMA chip PDMS membrane
Most importantly, for the first time, a system of experimental observation of the diaphragm on the side view under different frequency pulse and air force pressure was built In Figure 4-13(a~b), the experimental system involved the reverse microscope (Eclipse Ts2R, Nikon, Japan) connected with the high-speed camera (i-speed 3, Olympus, USA), the single-chamber PMMA/PDMS/PMMA device with dp@0 àm, and the external compressed air pressure source control The high-speed camera records the rapid movement of the PDMS membrane under the influence of various pneumatic pressures and different control signals (i.e., frequencies) The compressed air system consists of an air compressor that supplies compressed air to the microchamber and the high-speed valves (MHA2- MS1H, FESTO, USA) in which its operation could be modulated by a programmable WAGO controller (750-362, Germany) A software interface was developed in LABVIEW (National Instruments, Austin, Texas) for the control frequency of actuators Compressed air was routed to the solenoid valves, which were then connected via tubing to the fabricated chips containing functional elements A precise gas regulator was used to adjust the actuation of compressed air pressure The system was used for all experiments in this research The displacement of the PDMS membrane in the actuator diaphragm was observed and recorded by the high-speed camera to capture 500 frames/second The measured deflection of PDMS membrane in the single-chamber was investigated to vary air pressure from 50 kPa to 300 kPa, frequency from f=5 Hz to f@ Hz
Figure 4-13 (a) Experimental set up to observe the deflection of PDMS membranes (side view) under high-pressure and high-pulsation frequency (b) Enlarged image of the single microchamber corresponding to (a)
Figure 4-14(a) shows the top view of the deflection membrane (by using the optical view function of 3D laser profilometer machine) with the air pressure of 100 kPa, and Figure 4-14(b) shows the 3D profile of the deflection membrane corresponding to Figure 4-14(a) (by using the scanning function of 3D laser profilometer machine) As shown in Figure 4-14 (c), the center deflection and the radius of curvature of the deformed membrane were measured at various applied pressures, which vary from 50 kPa to 200 kPa The radius of curvature of the deformed membrane increases with increased applied pressure The membrane would rupture when subjected to higher pressures; however, the bond remained intact Following the results shown in Figure 4-11(a) and Figure 4-14(c), the curvature free-deflection of the PDMS membrane was always more significant than the designed PMMA spherical shape when the applied pressure was greater than 50 kPa Herein, we can confirm that under the static condition, no dead volume appears when PPMs are working at which actuated pressure of more than
50 kPa Please note that pneumatic force applied into the chamber could be
76 exceeded 700 kPa without any leakage or delamination of PMMA/PDMS/PMMA bonded
Figure 4-14 Deformed membranes under applied pressure: (a) Top view of the membrane at 100 kPa (b) Circular membranes at 100 kPa measuring by the 3D laser profilometer machine (c) Geometry of circular membranes at different applied pressures from 50 kPa to 200 kPa
Figure 4-15 Captured images of the deflection of PDMS membrane: (a) The magnitude and the frequency of pneumatic force are 50 kPa and 40 Hz, (b) The magnitude and the frequency of pneumatic force are 300 kPa and 40 Hz
Figure 4-15(a~b) shows the deflection of the PDMS membrane captured by the high-speed camera with the highest frequency f@ Hz (due to the smallest time resolution of the controller WAGO) at an applied pressure of 50 kPa and 300 kPa, respectively As shown in Figure 4-15(a), the PDMS membrane is not in contact with the PMMA chamber, which means it contains dead volume when PPMs are working at 50 kPa/40 Hz In contrast, in Figure 4-15(b), the PDMS is always in contact with the PMMA chamber even when conditions are activated at 300 kPa/40 Hz Then we used this method to measure the deflection of the PDMS membrane under various conditions
Figure 4-16(a) shows the relationship between the deflection of the PDMS membrane under different pulsation frequencies and pressures In general, deflection of PDMS membrane will be decreased when applied high pulsation frequency Significantly, deflection of the PDMS membrane rapidly decreased at the pressure of 50 kPa That means the dead volume will appear when PPMs operate at a high frequency When applied more amplitude of pressure of more
PDMS membrane PDMS membrane 200 àm
78 than 100 kPa, the PDMS membrane was always close entirely to the half- spherical PMMA with a frequency of less than 20 Hz However, the frequency applied was more than 20 Hz, and the PDMS membrane was not closed tight with the chamber
Figure 4-16 (a) Deflection of PDMS membrane under various frequencies and pressures (b) Respond time of PDMS when air force applied and released
A c tiv e d o f P D MS me mb ra n e ( ms )
As observed in experiments, the PDMS membrane returned to origin immediately after air force was released in the diaphragm actuator Figure 4-16(a) shows the time-responded deflection efficiency of the actuator diaphragm It is confirmed that pneumatic actuation with higher air pressure is preferable to lower pressure for the time-responded deflection efficiency of the actuator diaphragm
At 300 kPa/40 Hz, it only needs 4 ms for the PDMS membrane to contact the PMMA chamber; however, needing up to 20 ms to deflect at a positional of 300 àm at 50 kPa/40 Hz When operating under high frequency, almost the peristaltic micropump always faces lagging behavior of the PDMS membrane No experimental setup showed this behavior ever If the operating frequency is too high, the air chamber cannot be ultimately charged and discharged, and the pumping rate will not increase However, with our proposed technique, when the air chamber was depressurized, the PDMS membrane could recover to the original position within 2 ms even high frequency of 40 Hz operated Here we can confirm the proposed PPMs could be worked under high pressure without any lagging effect on the diaphragm Our study also guarantees the repeatability and reliability of the PDMS diaphragm working in the micropump without a mechanical failure and hysteresis, even high frequency actuation
Based on excellent bonding, we can apply more pressure over 300 kPa In Figure 4-15(b), it is clear that PDMS membrane always closed tight with PMMA event pulsation frequency applied of 40 Hz We can observe that the PDMS membrane return to origin position immediately after air force released in the
80 diaphragm actuator Figure 4-16(a) shows time-responded deflection efficiency of the actuator diaphragm It is confirmed that pneumatic actuation with higher air pressure is preferable to lower pressure for time-responded deflection efficiency of the actuator diaphragm With pressure of 300 kPa, only need 4ms to PDMS membrane contact the PMMA chamber, however, needing up to 20 ms to PDMS deflect at positional of 300 àm when pressure of 50 kPa applied Almost the peristaltic micropump when operate under high-frequency always face lagging behavior of the PDMS membrane However, no experimentally setup showed this behavior If the operating frequency is too high, the air chamber cannot be completely charged and discharged, and pumping rate will not increase However, with our proposed technique, as shown in Figure 4-16(b), when the air chamber was depressurized, the PDMS membrane could recover to the original position within 2 ms even high-frequency of 40 Hz operated Here we can confirm the proposed PPMs could be worked under high pressure without any lagging effect of the diaphragm
Our study also guarantees that the repeatability and reliability of the PDMS diaphragm working in the micropump without a mechanical failure and hysteresis even high frequency actuation, and please note that the reliability was excellent after using for more than one million actuation cycles Although we can increase depth of chambers to increase stroke volume, however, as shown here, high pressure operation is the key point those other methods cannot be achieved before
Extraction of blood plasma from human blood by integrated PPM
We characterized the performance of integrated inertia microfluidic device for blood plasma separation only in a single-device without any external pumps or valves The integrated-micropump consist of the PPM and a spiral microchannels with a trapezoidal cross-section area The PMMA mold was manufactured using the micro-milling technique which precisely milled a positive print of the spiral microchannels, as shown in Figure 4-23, then a PDMS spiral microchannel was made by casting on this mold The slanted four-circular-loops- spiral channels have a trapezoidal cross-section with a width of 500 μm, and inner and outer heights of 70 μm and 40 μm, respectively [63]
Figure 4-23 (a) The micromachined PMMA mold for a positive print of the spiral microchannels (b) Enlarged 3D image of a trapezoidal cross-section
The spiral PDMS channel was bonded with the micropump following the process of bonding mentioned above The flow rate was controlled by suitable frequency and air pressure to adjust the ranged of diluted blood 45× into the device, then blood cell concentration in the inner and outer outlet was observed
89 to evaluate the effectiveness of the integrated device Furthermore, a hemocytometer was used to quantify the blood cell concentrations in the collected plasma to examine the separation performance The separation efficiency is calculated by measuring the number of RBCs in the collected plasma sample and comparing the number to those in the original sample
Figure 4-24 (a) The preparation for blood plasma separation (b~c) SEM of trapezoidal microchannels corresponding section A-A’ and section B-B’, respectively
Following the concept of the spiral channels, the flow rate of 1.5 mL min -
1 (i.e., at 200 kPa/5 Hz) provides the best inertial focusing among working flow rate ranges Figure 4-24(a) shows the integrated microfluidic chip and the experiment setup of blood plasma separation, and Figure 4-24(b~c) shows the SEM of trapezoidal cross-sectionals view of spiral microchannels After using a
Reservoir of plasma (2) Spiral channel
Reservoir of blood cells (1) Reservoir of blood Triple-PPM
90 micropipette infused 200 àL of diluted blood in the reservoir, by controlling the frequency and air pressure, diluted blood was transported through the outlet of the micropump and then drive into the filtration channel
Figure 4-25 (a) Picture of samples collected from the inner outlet (No 1) for blood cells and the outer outlet (No 2) for blood plasma after the first round (b) Separation efficiencies of the diluted blood after 4 rounds of separations
Figure 4-25(a) shows the experiment results after the first round, in which the inner centrifuge tube (remark No 1) has the blood cells while the outer centrifuge tube (remark No 2) has the separated blood plasma The experiment results clearly show that this blood plasma separation can be realized on-chip with inertial microfluidics, and the on-chip micro pump can provide the sufficient flow rate Each droplet of blood plasma collected from the outside branch is clear; the blood cells collected from the inner branch are dark red, indicating a high concentration of blood cells that are predominantly blood cells Moreover, no phenomenon of RBCs logging or leakage of microchannels was observed Most importantly, the separation efficiencies were close to 97% after four rounds (Figure 4-25(b)) Note that the separation process takes less than 3 min with a tiny
Origin 1st 2nd 3rd 4th
Se p a ra ti o n e ff ici e n cy (% )
91 volume of diluted blood The most significant advantages offered by the device are high throughput, a reduced capacity, and the elimination of the requirement for any external valves or pumps
By on-chip without external pumps, we can extract plasma blood with many benefits: easy useful for micro àTAS or LOC, small volume, easy operation, less time-consuming
Conclusion, limitations, and recommendation
Conclusion
Many microfluidic applications could benefit from hybrid devices made of PDMS and PMMA; however, the existing bonding methods could not generate sufficient bonding strength and limit such applications This research adopted a systematical method, Taguchi method, to maximize the bonding strength between PDMS-PMMA microfluidic devices by optimizing the process parameters at room temperature, capable of achieving the highest ever bonding strength of 622 kPa (tensile strength of 3000 kPa) without the need for complex operations or sophisticated equipment From experiment results, the bonding strength was maximized using the following parameters: A1 (GPTMS concentration of 1%), B2 (APTES concentration of 3%), C1 (RF power of 6.8 W), D1 (exposure time of 60 s), and E3 (oxygen pressure of 800 mTorr) The excellent durability of the resulting bonding indicates that this bonding method is well suited to microfluidic systems requiring high operation pressure inside the microchannel (402 kPa herein), high throughput capability (highest flow rate of 120 mL min −1 herein), and long microchannel path needed to shrink into a small area (30 àm between adjacent channels herein)
With this reliable and sufficient bonding strength, we further developed PDMS-PMMA microfluidics with various applications:
• The microvalves are capable of rapidly and completely closing the microchannel or yielding a high steady flow rate under operation pressure ranging from 100 kPa to 700 kPa
• We developed the microvalve system, including eight microchannels and seven microvalves in high-density configurations, was manufactured, in which opening/closing microchannels in any order and any combination to enable the transport and sequential merging of various liquids can be realized
• We proposed a fabrication process of pneumatic peristaltic micropumps (PPMs) with high pumping rate by combining rigidly of PMMA and elasticity of PDMS Following the procedure to achieve excellent bonding strength of PDMS/PMMA, the PPMs could be operated under high pressure Then, the dynamic deflection test of PDMS membrane actuator under high pulsation frequency was performed for the first time We proved that the PPMs have excellent actuator diaphragm dynamic behavior without any mechanical fatigue or dead volume, achieving the highest pumping rate of 3500 àL/min
• The micropump was integrated with a spiral microchannel with trapezoidal cross-section area and used to rapidly extract plasma from human blood within 3 minutes and with a small blood volume of 200 μL, with separation efficiency up to 97% By on-chip without external pumps, we can extract plasma blood with many benefits: easy useful for micro TAS or LOC, small volume, easy operation, less time-consuming
Limitation of the dissertation
• In the pneumatic peristaltic pump, when the pump is working also so have backflow Since the first chamber was activated to begin the new cycles, the half volume of the chamber moves back, while half volume of the chamber moves forward, the micropump has backflow when the first chamber is activated However, the liquid transport is still stable and has laminar flow
• When separating blood plasma, the collected plasma from the outlet has light hemolysis, due to the PDMS membrane deflected by the air pressure And when diluted blood moves in the spiral microchannels
• In separating plasma from diluted blood, we only tested at 45× diluted blood; we have not tested in other conditions.
Recommendations
• Prevent backflow of the micropump by changing the operating sequence
• Improving separation efficiency of blood plasma from diluted blood by developing an integrated microfluidic system that can create new cycle separation
• Research to replace the air actuators with different actuators can make the pump system more compact, portable, and user-friendly
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