Highly Integrated Microfluidics Design

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Highly Integrated Microfluidics Design

For a listing of recent titles in the Artech House Integrated Microsystems Series, turn to the back of this book.

Highly Integrated Microfluidics Design Dan E. Angelescu

artechhouse.com

Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the U.S. Library of Congress.

British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library.

ISBN-13 978-1-159693-979-0

Cover design by Vicki Kane

© 2011 ARTECH HOUSE 685 Canton Street Norwood, MA 02062

All rights reserved. Printed and bound in the United States of America. No part of this book may be reproduced or utilized in any form or by any means, electronic or mechanical, including photocopying, recording, or by any information storage and retrieval system, without permission in writing from the publisher. All terms mentioned in this book that are known to be trademarks or service marks have been appropriately capitalized. Artech House cannot attest to the accuracy of this information. Use of a term in this book should not be regarded as affecting the validity of any trademark or service mark.

10 9 8 7 6 5 4 3 2 1

Contents Preface

ix

1

Microfabrication Techniques

1

1.1

Introduction

2

1.2 1.2.1 1.2.2 1.2.3 1.2.4 1.2.5

Manufacturing Processes PDMS Molding and Soft Lithography Metallic Fabrication Plastic Fabrication Paper Fabrication Silicon / Glass Fabrication

3 3 7 14 17 17

1.3 1.3.1 1.3.2 1.3.3 1.3.4

Technology Selection Criteria Required Functionality Physical Requirements Chemical Resistance Cost of Production

35 36 45 47 48

1.4 Conclusions References

49 50

2

Microfluidic Building Blocks

55

2.1

Introduction

56

2.2

Functional Approach

56

2.3 2.3.1

Single-Phase Fluid Manipulation Active Valves

57 58

v

vi


2.3.2 2.3.3 2.3.4

Passive Valves Pumping Techniques Single-Phase Mixing

71 75 91

2.4 2.4.1 2.4.2 2.4.3

Droplet-Based Fluid Manipulation Droplet and Emulsion Generators Droplet-Based Mixing Droplet Manipulations

99 101 107 108

2.5 2.5.1 2.5.2 2.5.3

Detection and Measurement Techniques Optical Measurements Chemical Measurements Physical Measurements

114 115 120 123


126 127

3

Microscale Physics

139

3.1

Introduction

140

3.2

Diffusion Laws

141

3.3 3.3.1 3.3.2 3.3.3 3.3.4 3.3.5 3.3.6

Fluid Dynamics Mass Conservation Equation Momentum Equation Fluid Forces Navier-Stokes Equation Stokes (Creeping) Flow Equivalence Between Fluidic and Electrical Circuits

143 144 145 147 151 154 166

3.4 3.4.1 3.4.2 3.4.3

Surface Tension and Wetting Surface Tension Forces Wetting Considerations Surfactants and Marangoni Stresses

172 173 174 176

3.5 3.5.1 3.5.2 3.5.3 3.5.4

Electrical Phenomena in Microfluidics Electrophoresis Dielectrophoresis Electro-Osmotic Flow Electrowetting (Electrocapillarity)

178 178 179 180 181


182 183

Contents

vii

4

Microfluidic Design

185

4.1

Introduction

186

4.2 4.2.1 4.2.2 4.2.3 4.2.4 4.2.5

Technology Functional Description Integration: Monolithic Versus Hybrid Material Choices In-House Versus Outsourced Fabrication Fabrication Process: Definition and Optimization

186 187 188 193 194 196

4.3 4.3.1 4.3.2

System Design and Optimization Mask Design Fluidic System Optimization

205 205 210


216 218

5

Integrated Microfluidic Systems

221

5.1

Introduction

222

5.2

Remarks About the Microfluidics Market

223

5.3 5.3.1 5.3.2 5.3.3 5.3.4 5.3.5 5.3.6 5.3.7

Selected Commercial Technology Examples Microfludic Large-Scale Integration: Fluidigm RainStorm Droplet Technology: RainDance Modular Plastic Microfluidics: ThinXXS Resonant Silicon Microtubes: ISSYS Nanopump MEMS Drug Infusion: Debiotech CMOSens Sensor/Electronics Integration: Sensirion Integrated Gas Chromatography Microsystem: C2V

225 225 229 232 235 237 240 242


244 246

About the Author

249

Index

251

Preface

ix

x

Motivation The field of microfluidics, until relatively recently considered an academic curiosity involving high manufacturing costs and having little application potential, is currently gaining acceptance enthusiastically in a wide range of industrial and academic fields. This rapid turnaround has been made possible, on one hand, by the development of novel fabrication, sensing, and fluid manipulation technologies, and on the other hand by important advances in microfluidic integration. Small-scale devices capable of performing complex microfluidic sample preparation and analysis functions now exist in many markets, and many more applications are actively being developed in academic and industrial research institutions. In these circumstances, it is an uneasy task to write a book about the continuously evolving frontline of microfluidics. I chose to concentrate on the theme of integration in microfluidic systems for several important reasons. On one hand, I believe that the integration aspects in microfluidics are inadequately covered in current literature. On the other hand, much of the current progress in microfluidics is coming from the capability to integrate multiple fluidic functions within an integrated system. The final reason, and perhaps the most important, is that the microfluidic community is currently fragmented, with different groups developing individual technological integration solutions adapted to their specific setup and environment. This lack of consensus is due to the complexity of microfluidic systems and to the extremely wide spectrum of possible applications. I hold the belief that unified approaches can be developed for dealing with the diversity encountered in microfluidics, and that such convergence of technological solutions will lead to faster and more efficient development of the field. While far from claiming to have found an answer, this book does, I hope, shed some light in the right direction.

Book Structure This book consists of five chapters. Chapter 1 provides a broad review of techniques currently being used for manufacturing microfluidic systems. It describes the different available technology solutions, with their relative merits, and provides a guide for selecting the right manufacturing method for the application at hand. Chapter 2 takes a functional approach to microfluidic design, breaking down the system into individual specialized building blocks (valves, pumps, sensing devices and so on) and describing the currently available solutions. Chapter 3 describes the main laws of physics that govern the behavior of fluids in microfluidic systems, providing the basic knowledge

Preface

xi

required to deal with the complexities of microfluidics (the focus is on breadth and not on depth, but references are provided for further in-depth reading). Chapter 4 attempts to describe the main tools that are required for designing, simulating, and modeling microfluidic devices. The focus here is on practical applications, and specific design examples are given along with examples of back-of-the-envelope calculations and simulation results. Chapter 5 provides an overview of several commercial microfluidic integration approaches, the choice of examples covering a wide variety of technologies.

Intended Audience This book is primarily intended for scientists and engineers involved in the design and development of complex microsystems that need to integrate multiple functionalities. Since the material is drawn from my experience in the field as a scientist involved in both industrial and academic research, I hope that this book will benefit both these communities. The main purpose of this book is to become a useful guide to readers who are starting to become involved with microfluidics, while acting as a (limited) reference to those well familiarized with the field. It is by no means comprehensive (which could have not been possible in a project of this size), but it does provide good references for a detailed reading on most topics covered. The text tries however to provide enough background material to be self-contained, thus being accessible to graduate and advanced undergraduate students learning about the field.

Acknowledgments I would like to thank several people who provided guidance and advice in writing this book, particularly Professor Howard Stone, Dr. Olivier Vancauwenberghe, Mr. Bruno Mercier, Ms. Katia Blanchard, and, last but not least my father, Dr. Nicolae Angelescu. I would like to thank Dr. Ram Shenoy for encouraging and guiding my early entrepreneurial activities. I would also like to acknowledge Mr. Mark Walsh at Artech House, who graciously provided the flexibility I required in writing.

1 Microfabrication Techniques

1

2


1.1 Introduction The advent of microfluidic technology has often been compared to that of microelectronics—the possibility of integrating multiple fluidic functions on a single device, in a manner similar to the integration of multiple electronic functions within an integrated circuit, is indeed very appealing in its potential to revolutionize many industries. The resulting lab-on-a-chip concept, where a single device performs multiple manipulations and analyses of fluidic samples, often with better accuracy and faster and cheaper than using traditional laboratory methods, has been demonstrated in several applications. The predicted “revolution,” however, is still waiting to happen: compared to integrated circuits, which had become ubiquitous within a few years of introduction in the late 1950s, microfluidics is still struggling to find a proper place in the global fluid device/sensor market. Despite a tremendous initial success in ink-jet printer head technology (which was developed more than 20 years ago, and is today one of the few mass-market applications of microfluidics), and the more recent development of several applications in the medical, biotechnology, and drug-discovery industries, microfluidics remains a niche technology with which most people do not interact in their day-to-day life. The electronics revolution was enabled by the development of silicon microfabrication, a universal manufacturing technique capable of producing electronic circuits of unprecedented complexity and capable of reproducing virtually all functions of traditional circuits. One of the principal reasons for the relatively slow market emergence of microfluidics is the lack of a similar universal fabrication technology, which at the same time is cheap, can be implemented in batch processes, and can accommodate all possible microfluidic functions: fluid pumping and control, thermal management, sensing of various parameters, optics, and so forth. Many different fabrication methods exist, each having their own strengths and weaknesses—the microfluidic system designer is ultimately responsible for selecting the manufacturing method best suited for a particular application, the choice often involving several trade-offs or use of hybrid approaches. This chapter will focus on fabrication technology, trying to achieve a triple goal: • Provide an extensive overview of available microfluidic and MEMS manufacturing processes; • Compare the characteristics of different manufacturing technologies, with a particular emphasis on integration capabilities;


3

• Guide the reader through the technology selection process for a specific application. This chapter is intended neither as a primer nor as a reference book on micromanufacturing technologies. A number of excellent reference books [1– 7] exist that review such technologies, and the reader is strongly encouraged to consult them for in-depth coverage on any particular aspect treated in this chapter. When new technological aspects are presented, which have been developed in recent years or only recently adapted to microfluidic fabrication, and which are insufficiently covered in existing reference texts, the original publications will be cited. For well-established technologies, however, which are thoroughly covered by existing books, specific citations will be omitted.

1.2 Manufacturing Processes 1.2.1 PDMS Molding and Soft Lithography Soft lithography designates a series of techniques involving the curing of an elastomer on a mold surface with a predefined topography (usually achieved by photolithographic means), followed by the separation of the elastomer from the mold. The resulting molded “soft” elastomeric part reproduces a negative of the mold topography. The part can subsequently be bonded to other surfaces, which may also be of elastomeric nature. This allows the fabrication of stacked structures, a feature that proves to have profound implications for the integration of various functionalities in an elastomeric device. Poly di-methyl siloxane (PDMS) is a transparent silicone elastomer, initially developed for encapsulation and potting applications in the electronics industry. Due to its optical transparency and ease of processing, it has become a material of choice for soft lithography. The capability to develop microfluidic prototypes from the concept phase to device fabrication, all within a short timeframe and without requiring much training or expensive software or equipment, has enabled a real revolution in the field of microfluidics: academic groups and companies all over the world are utilizing PDMS soft lithography in the most diverse microfluidic applications [8]. While the PDMS molding technology has certain limitations, making it less than ideal in many implementations, its role as a catalyst in developing microfluidics from a lab curiosity to a technology generating tremendous interest in multiple markets cannot be overestimated. Since PDMS soft lithography can easily be implemented in the average laboratory without requiring access to expensive fabrication equipment or clean room facilities, it has become a de facto fabrication standard in many academic settings. For this reason, we decided to include a detailed

4


Figure 1.1 Typical PDMS molding process. Top left: The SU-8 master is prepared by photolithography, to provide a negative of the desired surface topography. Top right: The PDMS is poured (a surface treatment of the wafer may be required to prevent adhesion) and cured. Bottom left: The cured PDMS stamp is separated from the master template. Bottom right: Holes are poked for the fluid inlet and outlet tubes, and the PDMS is treated to improve adhesion and is bonded to either a glass slide or another PDMS substrate.

soft lithography process description in this chapter. The typical design and fabrication process involves several steps. While alternate fabrication sequences exist, the following list, along with Figure 1.1, captures the most typical fabrication flowchart: • Mask design is achieved using a CAD software package. Typically, the low-resolution requirements of microfluidics allow the use of inexpensive plastic transparency masks. These can be ordered (and often received overnight) from commercial printing service companies, which often offer printing processes with minimum feature sizes down to 20 µm; most consumer printers are not capable of printing transparencies with enough contrast and resolution for any realistic application beyond the most basic designs. For higher resolutions, chrome quartz masks can be used.


5

• The mold is manufactured using photolithographic techniques. Photoresists that can be cast in thick layers while providing vertical straight walls are preferred—a typical choice involves a negative resist from the SU-8 family, but other choices with similar capabilities exist. The casting of the photoresist is typically achieved using a spincoating process, with the final thickness being controlled by the spin speed and the viscosity of the particular resist formulation chosen. For the best spin-coating results, the selected resist formulation should form a film of the desired thickness at a spin speed ideally between 1,000 rpm and 4,000 rpm. Heat curing of the photoresist prior to and after exposure is required for the negative resists which are most commonly used for this step. Heat curing of thick resists requires special attention to minimize thermal stresses that may lead to cracking of the photoresist layer—in particular, very gentle temperature ramps of a few degrees per minute should be used. The usual substrate material for the mold is a silicon wafer, which can be purchased inexpensively (300 to 700 µm thick, arbitrary orientation, test grade, single-side polish, arbitrary oxide thickness). • Additional photoresist layers may be cast and exposed to create structures having more complicated topographies. • The mold may be coated with a release agent, to facilitate the separation of the elastomer after curing. The coating is typically performed from a vapor phase; the release agent molecules bond to the substrate material, providing a low-energy, nonstick surface—one of the more common release agents used to coat silicon wafers is trimethyl chlorosilane (TMCS). • The elastomer formulation is prepared. This involves mixing a base elastomer with a curing agent that will promote cross-linking (both components typically come in liquid form). Single-component elastomers exist; these typically cure by exposure to moisture from the air or to UV light. • The elastomer mixture is mixed and degassed. Mixing may be achieved manually, or, alternatively, an automated mixer can be used. The degassing step is required to avoid trapping air bubbles in the cured elastomer, which may otherwise compromise the mechanical properties and optical clarity of the final device. It is typically achieved in a vacuum chamber, under a relatively low vacuum (roughing pump vacuum being sufficient). As gas bubbles in the liquid elastomer mixture expand, they raise to the surface

6


of the elastomer, creating a foam—the vacuum level may require adjustments to control the foaming process. • The mold wafer is placed in a recipient, and covered with a layer of the degassed elastomer. • An additional degassing step is required to eliminate gas bubbles that may have formed during the pouring process or may be trapped in the corners of the mold microstructure. • The curing is initiated. This may involve curing inside an oven (curing for a couple hours at 60o C to 80o C is typically required for PDMS), or exposure to another curing agent (moisture, UV light). The curing process will define the degree of cross-linking in the material, and hence the mechanical properties of the final device: longer or higher temperature curing typically results in stiffer devices. • The separation of the elastomer from the mold wafer is performed; cutting of the elastomer to the desired shape may be achieved using a razor blade before or after the separation. Care needs to be taken not to shatter the mold when using fragile substrate materials such as silicon or glass wafers. • Fluidic connection holes are made through the elastomer, for example, by using a blunt syringe needle with the outer diameter similar to or slightly smaller than that of the tubing that will be used for fluidic connectivity. The needle used for making the ports may be left in the elastomer to preserve the shape of the ports during subsequent processing steps. • The elastomeric surfaces may be activated using either an O2 plasma or exposure to ozone from a UV-ozone cleaner. The plasma activation process is highly dependent on the specific equipment and on environmental conditions such as humidity levels. It needs to be customized for the particular setup. This step is required for subsequent bonding of the elastomer to a different substrate, which may also be elastomeric in nature. Elastomeric pads can also be easily bonded to glass and oxide-covered silicon wafers, which confers rigidity to the device but creates channels having different wetting conditions on different channel walls, which may be an important issue if the device is to be used with multiphase flows. Preferably, elastomeric pads are bonded to a glass slide covered with a thin elastomeric layer that was spin-coated and then cured. The glass


7

slide itself may be patterned, most often with electrodes for applying electric fields in the final device. • The different surfaces to bond are aligned, either visually under a microscope, or using alignment fixtures such as pins or microfabricated features. This step needs to be performed immediately after the plasma oxidation of the surfaces, typically within 1 minute. The surfaces are then brought in contact and cured together to enhance the bonding. • Tubing is inserted in the preformed fluidic connection holes, to allow fluidic connectivity and/or pneumatic or hydraulic actuation. The tubing is typically held in place by friction with the elastomer material; however, more exotic holding fixtures may be required to enable operation at higher pressures. These may include metallic fixtures that are held together using bolts and which can press on the elastomer to prevent delamination at the bonded interface (this, however, will not prevent the deformation of channels due to internal pressure). The same holder can be machined to accept, position, and hold the tubing, limiting the possibility for the tubes to be pushed out of the holes by the system pressure. • Additional fabrication steps may be required in certain cases to provide additional functionality to the device. For example, additional bonding steps may be performed to create a multilayered device, or optical fibers may be inserted and sealed within channels to allow optical connectivity.

1.2.2 Metallic Fabrication Manufacturing microfluidic devices out of metallic materials is possible; however, the costs of fabrication per device are typically higher than for any other techniques listed here. This fabrication method is justifiable in some cases, for example, when the part manufactured is a mold or stamp subsequently used to fabricate multiple replicas out of a different material (e.g., in an injection molding process). Other cases where metal fabrication may become necessary is in very high pressure and/or high temperature (HPHT) applications, in environments subject to severe shocks and vibrations, or when chemical resistance requirements eliminate other choices of materials.

8


Figure 1.2 Complex metallic microparts manufactured using high-speed conventional machining. The smallest scale machined here is on the order of 50 micrometers. (Used with permission from Atometric, Inc., www.atometric.com.)

1.2.2.1

High-Speed Conventional Micromachining

Conventional drilling, milling, and lathe operations can be miniaturized to produce features down to the micrometer scale. The miniaturization push coming from several industries has fueled a number of efforts in this direction. Tabletop milling machines have been developed that are capable of turning at speeds of tens and even hundreds of thousands of RPM, while allowing 5-axis control of the part orientation and providing translation stages with submicron accuracy. Tooling has also responded to the miniaturization push, with several manufacturers offering end-mills down to 25 µm in diameter. Unfortunately, not all manufacturing operations are amenable to highspeed machining, and often multiple premachining or grinding steps may be required to prepare the part and an appropriate fixture for the final micromachining process. Burring and chipping are important concerns: some metals may develop burrs that are difficult to remove during highspeed machining, and metal chips may become trapped in the manufactured microstructure. Contamination from cooling fluids may be an issue; however, many machines can be adapted for air cooling. Despite its deceiving name, the machine time required to manufacture even relatively simple designs using high-speed micromachining may actually be quite long: while linear accelerations and machining speeds can be very high, the cutting depth needs to be maintained relatively small to avoid tool overheating and disintegration, multiple steps at different depths thus being necessary. Nevertheless, conventional metal micromachining may be the only fabrication technique capable of creating certain part geometries (Figure 1.2), particularly when complex nonplanar structures are required. In addition, high-speed micromachining is not limited to metals only: plastic and ceramic materials


9

Figure 1.3 A micro-EDM computer-controlled machining unit performing the final machining operations on a turbine blade. The microelectrode is brought in close proximity of the surface, both being immersed in a jet of dielectric fluid. Besides its electrical role, the fluid also acts as a coolant and carries away debris from the surface. (Used with permission from Leer Technologies, www.leertech.com.) )

can also be machined using this technique, making it a versatile prototyping method. 1.2.2.2 Micro-EDM

Electro discharge machining (EDM) is a fabrication technology for electrically conductive parts based on material removal due to high-energy discharge between an electrode and the machined part [9]. Typically, the part is submerged under a jet or in a bath of dielectric fluid, and the electrode moves near the surface of the part (without actually making physical contact); Figure 1.3. Very small electrodes can currently be manufactured, capable of machining features down to a few tens of micrometers—this capability has launched the field of micro-EDM manufacturing [10]. Currently, machines exist that are capable of submicron part positioning as well as 5-axis control. Micro-EDM processes are capable of producing high aspect ratio features in a part, with no burrs and little contamination. EDM can machine almost any conductive material, including very hard alloys used as injection-mold materials. The micro-EDM technology has several important limitations: material choice is limited by the conductivity requirement and the process may lead to the contamination of the machined part by the dielectric

10


fluid used or by electrode wear—subsequent solvent cleaning and chemical treatments may be required to remove contamination. EDM technology also shares some of the disadvantages of high-speed machining, particularly related to the fabrication speed and the need of premachining steps. A related technique is wire-EDM, where a conductive wire is used to cut vertical slots through a metallic layer. The wire-EDM technique allows very narrow, high aspect ratio through channels to be machined in a plane metallic part.

1.2.2.3

LIGA and Derived Techniques

LIGA is a German acronym for “LIthographie, Galvanoformung und Abformung,” meaning lithography, electroforming, and molding. As the name implies, the method was originally developed in Germany, at the Karlsruhe Nuclear Research Center. In its original form [11], the LIGA method involves a first step of deep X-ray lithography using a highly collimated synchrotron radiation, during which a mask pattern is transferred into a thick film of X-ray sensitive photoresist, such as PMMA, cast on a conductive substrate (or base plate). The photoresist is subsequently developed, and the resulting microstructure is placed in an electroplating bath, allowing metal plating in the areas where the photoresist had been removed during development. The resulting structure may be used by itself, it may become part of a mold insert for plastic injection molding, or it may be subjected to additional LIGA steps to create multilevel structures. Multilevel LIGA has a number of additional processing requirements, namely, the planarization of the electroplated layers, the adhesion of a new resist film to the previously planarized layers, and the alignment of the additional lithographic masks with the previously built structures. Variants of the technique have been developed that eliminate the constraint and fear factor of needing synchrotron radiation for the exposure of the photoresist layer. In particular, UV-LIGA takes advantage of the relatively recent development of thick UV-curable photoresists (such as the SU-8 family of resists), which allow high aspect ratio structures to be created by standard photolithographic techniques. LIGA suffers from a number of disadvantages: complicated and expensive processing, limitation to two dimensions (or to several stacked planar layers for multilevel devices), and limited materials choices. It is, however, capable of performing batch fabrication of metallic microparts, of generating metallic structures with very high feature density and intricate planar geometries (Figure 1.4), and of producing high-quality micromolds for the plastic injection molding of microfluidic parts.


11

Figure 1.4 Structure fabricated using LIGA technology. Extremely high aspect ratios can be achieved using this technology, due to the excellent collimation of synchrotron radiation. (Used with permission from HT Micro, www.htmicro.com.)

1.2.2.4 Photoetching

Photoetching is a relatively inexpensive method of batch-producing flat metallic parts structured at the microscale. The method consists of coating both sides of a metal film with a suitable photoresist, exposing the resist through a lithographic mask, developing it, and then placing the film in a metal etchant bath. The photoresist may be patterned on one side only, or on both sides (in which case back-side alignment is necessary during the lithographic step). By carefully controlling the depth of the wet etching step, channels of controlled width can be etched completely through the film or only partway. The side-wall profiles of etched channels are typically curved, a caveat of the isotropic wet etching method used. Very complicated planar geometries can be realized (Figure 1.5); three-dimensional geometries can be constructed by bonding, laminating, or simply clamping several photo-etched layers together in a sandwich-like structure, or by bonding etched layers to metal parts manufactured using other machining techniques. A wide range of metallic materials are suitable for photoetching (stainless steels, gold, titanium, nickel, and so on), as well as some polymeric materials (e.g., polyimide). 1.2.2.5 Laser Ablation

The laser ablation technique consists in focusing very short laser pulses (on the order of several picoseconds to several nanoseconds) to a small area of a substrate. The material in a very localized area is heated and effectively vaporized, creating a small crater. The lateral size of the crater is controlled

12


Figure 1.5 The photo etching of thin metal films allows low-cost batch production of metallic parts. The film can be through-etched or only partially etched. Subsequent assembly, alignment, and bonding of such parts can result in complex metallic microfluidic devices. (Used with permission from Microphoto, Inc., www.microphoto.net.)

primarily by the laser focusing spot size, which may reach submicron dimensions, whereas the depth of the ablated region is primarily controlled by laser power and pulse duration, which together define the amount of energy packed in one pulse [12]. The final depth can thus be tweaked quite precisely by adjusting the number of pulses, which gives laser machining the intrinsic capability to manufacture three-dimensional structures. By moving the laser beam around, continuous microchannels may be generated (Figure 1.6), as well as dense arrays of holes. Laser machines with five control axes exist nowadays, enabling very complex laser micromanufacturing. Laser machining is a very slow process, however, and therefore costly. It is not amenable to a high-volume production of microfluidic parts, but it can be an excellent prototyping tool for R&D-type activities. Lasers can be used for machining different exotic materials, including many semiconductors, glass, ceramics, plastics, and metals; lasers can also be used for bonding and sealing plastic devices. 1.2.2.6

Metal Sealing and Bonding Techniques

Fabricating a complete microfluidic device typically requires bonding several parts to create a completely enclosed channel or other sealed geometries. Several methods for joining similar or dissimilar metals exist, as well as for bonding metals to other materials such as glass. The main techniques applicable to building metallic microsystems involve brazing and different forms of welding.


13

Figure 1.6 Laser machined microchannels in PMMA plastic. The technique can be used to machine other materials, such as glass, metals, and ceramics. (Reproduced with permission from [13].)

One of the oldest and most widely utilized technique for joining metals is brazing, where a brazing filler material (also called a brazing alloy) with a melting point lower than either metal to be joined is applied near the joint, and the parts are heated, typically in an oven, above the filler’s melting temperature. The filler is then pulled by capillarity into the joints, and the resulting part is cooled to resolidify the filler and complete the bonding process. The parts need to be capable to withstand uniform heating at relatively high temperatures, typically 100o C or more above the filler’s melting point (in many cases, this precludes the assembly of some hybrid microsystems containing low melting-temperature materials). Often, precleaning of the parts is required (e.g., to remove contamination and native oxide layers that form on most materials). Such treatments may include solvent cleaning steps, the use of a flux material with the brazing filler, or exposure to a controlled atmosphere (e.g., partial hydrogen atmosphere or vacuum), prior to and during the brazing [14]. Parts must also be machined with very small clearances (of a few micrometers) to allow for the capillary filling of the gaps. Since machined microchannels may be comparable in size with the machining tolerances of the parts, the channels risk being clogged with filler compound during the brazing process. This limitation of brazing can be overcome by holding tight machining tolerances, employing appropriate holding fixtures, and using very thin layers of brazing filler (on the order of a few microns), preformed on the surfaces to be bonded, for example, by a plating process (nickel, copper, silver, and gold are commonly used). A good brazing joint can have outstanding strength: it is possible, under failure testing, that base materials will tear before the brazing interface delaminates. Welding is a generic name that covers several methods of joining similar metals by subjecting them to heat treatments without employing a dissimilar

14


metallic alloy interlayer or filler material. Welding may involve melting the base materials to be bonded (such as in arc, laser, or e-beam welding), or it may be performed by applying high pressure at temperatures slightly below the melting point, such as in solid-state (diffusion) welding. Most of the welding techniques that are amenable to joining microfabricated parts do not involve the addition of material (like the welding rod utilized in TIG welding). Microfabricated assemblies also require high precision welds that can be automated, thus precluding manual welding. The three welding processes most suitable for metallic microfluidic fabrication are e-beam, laser and solidstate welding. In e-beam and laser welding, the energy required to locally heat the base materials above their melting temperature is provided by an accelerated electron beam, or a focused laser beam, respectively. In solid-state welding, by contrast, the entire assembly to be bonded is heated to a temperature a few tens to hundreds of degrees below the melting temperature, and pressure is applied to keep the interfaces to be welded in intimate contact. The enhanced atomic diffusion occurring at high temperature results in metallic grains extending across the interface. To achieve good bonding by this technique, the parts to be joined must be very planar, smooth, and clean. Organic materials and native oxides, in particular, need to be eliminated from the interface. This is often achieved by solvent- and acidbath precleaning, and by performing the bonding process in a controlled atmosphere of hydrogen and inert gas. Bonded parts can be machined by conventional processes after the bonding. This bonding method is particularly well suited for laminating multiple thin metallic parts [15], machined by any of the techniques described in the preceding sections. While a relatively slow process (a typical diffusion-bonding run, with all the preparation steps, may last several hours), solid-state welding is amenable to batch fabrication: multiple parts can be bonded in a single run, either by stacking or by laying them side by side in the diffusion-welding oven. 1.2.3 Plastic Fabrication 1.2.3.1

Injection Molding

Many commercial microfluidic devices are fabricated out of polymeric materials using plastic injection molding. In this method, the plastic is heated above its glass transition temperature (or its melting point, in the case of a crystalline polymer), and then injected (sometimes under very high pressures) into a multiple-part metallic mold. The assembly is then cooled, the mold is opened, and the part is separated from the mold. Injection molding can be very fast - fabrication of one part may take as little as a few seconds.


15

Figure 1.7 High aspect ratio plastic channels obtained using hot embossing using an etched silicon wafer as a tool. The channels are approximately 2 µm wide. (Reproduced with permission from [16]. Copyright Wiley–VCH Verlag GmbH & Co. KGaA.)

The process imposes some limitations in geometry of the part; relatively complex parts can, however, be achieved, with intricate three-dimensional geometries, including threads, and features down to the micrometer level. Many types of thermoplastic materials can be micromolded—cyclic olefin copolymer (COC), polymethylmethacrylate (PMMA), polycarbonate (PC), polypropylene (PP), and polystyrene (PS) are commonly used. Micromolding can also be used with some high-performance engineered plastics such as polyetheretherketone (PEEK) and Ultem, materials with extraordinary mechanical strength, and chemical and temperature resistance that are commonly used in the medical, oil, and aerospace industries. The micromold fixture is typically fabricated using one of the methods described in Section 1.2.2. Very strong metals should be used for mold fabrication to avoid degradation over thousands, possibly millions, of injection cycles. 1.2.3.2 Hot Embossing

Embossing is another method commonly used for surface plastic microfabrication. It involves pressing a microstructured material into a preformed plastic part at a temperature close to the glass transition of the polymer, typically under a vacuum to avoid defects from gas bubbles trapped in the microstructure (Figure 1.7). The process involves a relatively long per-part cycle and is limited to planar geometries, but its simplicity and low setup cost make it the process of choice in many micromanufacturing applications. Embossing tools made of metal (Section 1.2.2), silicon (Section 1.2.5), and even epoxy have been successfully used. Microfluidic devices with

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Figure 1.8 Left: Bonding of plastic microfluidic components using heat and pressure—the temperature in the hot press or oven needs to be close to the melting temperature, and sufficient pressure is required to bring the two surfaces in conformal contact. Right: Thermal bonding using laser heating requires one substrate to be transparent and the other absorbent at the laser wavelength. The heating is applied locally, at the interface between the two substrates.

completely enclosed channels can be fabricated by bonding a top layer to the device; fabrication using many plastics can be achieved this way, including high-performance materials like PEEK [17]. 1.2.3.3

Plastic Sealing / Bonding Techniques

A large number of techniques exist for bonding plastic layers [16]. The easiest and most universal method is to use adhesive bonding. This provides the capability to bond virtually any materials, provided an adequate adhesive is identified. Care must be taken to avoid the wicking of adhesive within the microfluidic channels, which could cause clogging. Adhesive bonding can be used in some applications, particularly where the microfluidic geometry is simple, the size of the channels is large, and there is no need to separate nearby channels. Similar to adhesive bonding, lamination involves the adhesion of two plastic films (one of which is precast with a thin adhesive layer a few microns thick) using a hot-rolling process. During the rolling process the adhesive melts and then solidifies again when it cools down, thus bonding the substrates. Just as in adhesive bonding approaches, capillary wicking of the molten adhesive can clog small microfluidic channels; however, this is less likely to happen due to the limited amount of adhesive available within the precast film. Solvent-mediated bonding approaches can also be used to bond plastic surfaces, whereas one or both surfaces are exposed to an adequate solvent, which provides mobility to the polymer molecules within the plastic. By


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bringing the two surfaces into contact, the polymer molecules become intertwined, a process which results in a permanent bond once the solvent diffuses away. The same process can be used to bond a plastic channel plate to a glass substrate—the channel plate is pressed against the substrate, and the solvent is applied to the interface and sucked within the plastic-glass gap by capillarity. This method, however, does not allow complex channel structures to be built. Thermal methods (Figure 1.8) are the most common techniques for bonding plastic chips. In the classical bonding process, the two plastic parts are brought into contact in a hot press, and heated close to the melting temperature. The increased mobility of the polymer chains results in enhanced diffusion across the interface, effectively bonding the chips. The pressure and temperature need to be carefully controlled, or else the devices will be damaged during the process. The heat for thermal bonding methods can also be provided locally by a laser (in laser welding approaches), which allows significantly more complicated geometries to be built. 1.2.4 Paper Fabrication An original fabrication method that uses low-cost paper and double-sided adhesive tape for implementing microfluidic conduits has recently been developed [18] (Figure 1.9). The approach is based on patterning hydrophobic regions onto hydrophilic paper using a photolithographic technique, which creates “channels” on the paper where the fluid is drawn by capillarity. Complex three-dimensional topologies are possible by creating connections between different layers. These are made by bonding paper layers together with double-sided tape in which holes have been machined by laser-cutting (to allow liquid wicking through the holes between different paper layers, these are prefilled with a hydrophilic paste). This manufacturing technology, while certainly limited in capability compared to other techniques described in this chapter, has an enormous advantage: it involves low-cost materials and low-cost fabrication techniques, which could lower total manufacturing cost enough to enable applications for single-use medical devices and other cost-sensitive domains where other technologies are prohibitive. 1.2.5 Silicon / Glass Fabrication Silicon is a very versatile material as a substrate for microelectronics, a substrate or structural part in MEMS devices, or both. A large number of processes and recipes (both mechanical and chemical) for processing silicon

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Figure 1.9 Low-cost paper-based fabrication technique, using capillary wicking to drive fluid in predefined channels patterned on a paper layer, and punched double-sided tape to make fluidic connections between adjacent paper layers. (Reproduced from [18] by permission of The Royal Society of Chemistry.)

have been developed over the past half century, giving the microsystems designer a wide range of options from which to choose. The material is abundant in nature, is relatively inexpensive, and can be readily purchased from numerous suppliers in wafer form, with well-controlled crystal orientation, thickness, and surface roughness. A related substrate material that is becoming increasingly popular in the MEMS world is the silicon-on-insulator (SOI) wafer, consisting of two silicon layers of well-characterized thickness separated by a thin layer of silicon oxide. One of the silicon layers is typically thin (from a fraction of a micron, to a few tens of microns thick), and is called the device layer, whereas the other layer is thicker and provides structural strength (the handle layer). Glass is one of the oldest materials discovered by man. The optical and mechanical properties of this material make it as interesting to the microsystems designer today as to the antique societies who accidentally discovered it millennia ago. It is a transparent, hard, and brittle material that is cheap to produce and recycle. Many types of glass exist, the main difference consisting of the types and concentrations of ions, oxides, and impurities present in the material. Glass containing boron oxide as one of the main ingredients is commonly called borosilicate or Pyrex glass. This type of glass has low thermal expansion coefficient (approaching that of silicon) and generally provides good optical transmission throughout the visible, near IR, and long-wave UV range. Most importantly, this type of glass is rich in sodium ions and can be bonded to silicon using anodic bonding without inducing significant thermal stresses, thus providing a good way to encapsulate MEMS devices and/or create fully contained microfluidic systems.


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The combined use of silicon and glass provides the most versatile fabrication technique for fluidic microsystems, with the potential of creating fully integrated systems containing fluidic passages, sensor elements, optical elements, actuators, and embedded electronics. Silicon and glass fabrication require the use of a clean room, making this technology relatively expensive and not within the reach of every academic laboratory. In addition, the clean room processing of silicon and glass implies a steep learning curve and relatively long fabrication times with little flexibility in changing the fabrication flowchart midway, making it less suitable for rapid prototype development compared to other techniques. These disadvantages have initially hindered microfluidics from turning into the popular research field that it has become today, through the advent of less demanding fabrication techniques such as soft lithography. Nevertheless, silicon/glass fabrication remains a technology of choice when high levels of integration are required, and is used in numerous commercial applications. Fabrication processes for both glass and silicon can be divided into lithographic (patterning of material), additive (deposition of material), subtractive (removal of material), and formative (a process through which the properties of a substrate or a previously deposited material are modified). Depending on whether the active device elements are manufactured by applying and/or selectively removing layers on the surface of the bulk silicon wafer or by machining the bulk silicon itself, silicon micromachining processes can be further divided into surface and bulk micromachining (this is particularly relevant to subtractive processes). Bonding between glass-glass, siliconsilicon, and glass-silicon wafers are treated as a separate class of fabrication processes. 1.2.5.1 Lithographic Processes

Lithography will be defined in the following as the process of transferring a pattern from one medium (called a lithographic mask) to another (usually a physical layer). Several processes can be described as lithographic (a summary is given in Table 1.1); however, in microfabrication photolithography is the accepted standard. Silicon and glass processing normally requires multiple photolithographic steps on one or both sides of the substrate. Alignment between subsequent lithographic processes is usually necessary; tolerances close to the micron level can be achieved by using specialized equipment (a mask aligner) and by designing specific alignment features on the mask. Double-sided alignment is required when both sides of a wafer need to be patterned.

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Table 1.1 Relative Comparison of Different Lithography Types Lithography Type Photolithography

Common Uses Conventional Si/glass processing Metal photoetch fabrication Soft lithography mold

Gray-scale

3-D fabrication Microoptics Complicated topography Incompatibility with photolithography

Shadow mask

Microcontact E-beam

Laser writing

Surface chemical patterning High-resolution patterns Photomask fabrication Alignment with random features Photomask fabrication Alignment with random features

Advantages Versatile Fast, batch-compatible Standardized process Inexpensive Versatile Inexpensive Simple processing Limited to depositions Inexpensive High-resolution capability Few nm feature size Direct from CAD

Disadvantages Feature size limited by λ Requires flat substrate Multiple chemical steps

Direct from CAD

Slow Expensive

Difficult to control Pattern limitations Low resolution Difficult alignment Poor alignment/registration Slow Expensive

• Photolithography The most common type of lithography is photolithography, by which a thin layer of photosensitive polymer is exposed to light (typically in the 254 nm UV range) modulated using a binary optical mask having transparent and opaque regions. Sometimes a gray-scale mask is used, allowing intermediate UV transmittance levels—the depth of exposure in a positive photoresist will depend on the exposure dose, and the postdevelopment thickness can thus be modulated, allowing three-dimensional surfaces to be “written” in the photoresist [19]. A similar effect may be achieved by using photoresist reflow [20] after the lithographic step as a way to create smooth thickness gradients. • Shadow-Mask Lithography Another type of lithography is shadowmask lithography, where a thin layer of material is deposited through a perforated structure (typically made of silicon or metal) onto a substrate. This type of lithography is used when the substrate has difficult topography, making photoresist casting difficult, or when the structure being fabricated cannot tolerate the normal photolithographic process steps. It has a number of disadvantages compared to photolithography, such as alignment difficulty, feature size limitations, and pattern topology constraints due to the nature of the shadow mask; a continuous aperture cannot, for example, fully surround a solid region. • Microcontact Printing and Nanoimprint Lithography Microcontact printing using soft lithography is another method of chemically patterning a surface at the micro- and nanoscales [21]; however, due to registration issues rising form the soft nature of the stamp, it cannot readily be applied to batch device fabrication. It can be


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used to transfer a high-density pattern, which may be beyond the capabilities of photolithography, to specific areas of the wafer. A similar technique is nanoimprinting, which relies on embossing a thin polymeric layer with a master having the negative topography, allowing reproduction of the pattern to scales as low as a few nanometers. This technique relies on hard materials and thus has the capability of good registration over the full wafer surface. It has been used to fabricate a variety of devices, from nanoscale photodetectors to quantum wires and dots [22]. • E-Beam and Laser Writing All types of lithography described above are parallel in nature, all areas of the substrate being simultaneously processed. Other patterning methods, serial in nature, exist. These are typically required for transferring a pattern directly from a CAD program to a substrate, and are used, for example, to create the optical lithography masks utilized in photolithography, or to directly create patterns that cannot be easily achieved other types of lithography (e.g., when very high feature densities and small feature sizes are required). Such CAD-to-substrate patterning techniques can also be used when alignment with random features preexisting on a substrate is required (e.g., when contacts need to be placed at the ends of a nanotube lying at a random position on a substrate [23]). Examples are e-beam writing and laser writing. The main disadvantages of such techniques stem from the serial nature of the writing, resulting in slow and expensive processes. 1.2.5.2 Additive Processes

• Spin Coating Spin coating is the method of choice for making thin polymeric films. The silicon wafer is placed on a chuck coupled to a spinner motor. The polymer or photoresist, dissolved in or thinned by an appropriate solvent, is poured onto the substrate. The motor is operated, causing the wafer to spin (ideally at 1,000–5,000 RPM, for 30 seconds or longer) and the polymer solution to spread in a uniform layer onto the substrate. Depending on the type of polymer, a baking step may be required after spin-coating to assure that all solvent has evaporated. The final thickness and uniformity of the layer are determined by a combination of several process parameters: polymer solution viscosity, solvent volatility, spin speed, and acceleration (ramp), spin time, and specific thermal processing. Polymer films ranging in thickness from few nanometers to hundreds of micrometers

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can be deposited this way. Other types of materials (such as spinon glass) can also be deposited this way, starting from a liquid-solid suspension. • Evaporation Evaporation is a physical vapor deposition process that relies on a target material being heated to a temperature close to its melting temperature, at which point target material atoms evaporate, travel ballistically through a high-vacuum chamber, and then deposit, or condense, on the substrate to be coated. The two most common heating methods are thermal (the material to be melted being placed in a crucible surrounded by a heating filament), and electron beam (the material, typically in the form of an ingot, being bombarded by electrons from a high-power electron gun). A large number of materials can be evaporated, such as most inorganic compounds, metals, oxides, and carbides. Noble metals (Au, Pt) do not adhere well to most substrates, and therefore need to be deposited on top of a thin metallic adhesion layer (typically Cr or Ti). Special attention needs to be given to the fact that some compounds may dissociate during the evaporation process, leading to a film composition different from the target. The presence of reactive gases even in trace amounts can lead to chemical reactions in the deposition chamber, resulting in film contamination. While usually this is an undesired side effect of an insufficient vacuum during the deposition process, it can also be a feature, allowing the deposition of certain chemical compounds (e.g., deposition of carbides by reactive evaporation in the presence of acetylene vapors). The temperature of the substrate may rise due to radiative heating from the source; however, evaporation is usually suitable for deposition on most plastic films. Due to the ballistic transport of the atoms through the evaporation chamber, this deposition method is prone to shadowing and does not achieve very good step coverage. This can also be a feature, facilitating subsequent liftoff processes. The ballistic transport from a point source to a large wafer can also introduce some thickness nonuniformity from purely geometric reasons, which can be minimized by placing the wafers oriented toward and relatively far from the target. • Sputtering Sputtering is a physical vapor deposition method that relies on atoms from a target material being ejected due to collisions with highly energetic gas ions, and then redeposited on the substrate to be coated. The plasma gas used in sputter deposition is typically inert (e.g., Argon), and is accelerated into the target by means of an electric field (DC or radio-frequency AC). One variant of sputtering


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deposition, utilized in the deposition of some oxide and nitride films, is reactive sputtering, where a reactive gas (O2 , N2 ) is introduced into the deposition chamber alongside the inert gas. Sputtering is a very versatile process, having a wide range of control parameters that dictate the film thickness, density, microstructure, and purity; this same feature also makes it a delicate process, requiring very careful control to assure reproducibility. It can be used for the deposition of a wide range of materials: precious metals, alloys, oxides, nitrides, carbides, borides, and semiconductors. Similar to evaporation, the sputtering of noble metals needs to be preceded by the deposition of a thin adhesion layer to assure good bonding to most substrates. Sputtering can be performed with the substrate at ambient temperature, making it a suitable process for deposition onto polymeric substrates. When the size of the target is larger than the substrate, sputtering offers good step coverage due to the random direction of incoming atoms. The films obtained by sputtering may be internally stressed, however. The stress can be minimized by adjusting process parameters, most notably the gas pressure or the substrate temperature during deposition. • (Plasma-Enhanced) Chemical Vapor Deposition Chemical vapor deposition (CVD) is a thin film deposition method relying on chemical reactions between gaseous precursors occurring on or near the substrate surface. CVD is usually performed in a CVD furnace at very high temperatures (as high as 1,000o C for polysilicon), which many materials cannot tolerate. CVD should therefore be performed as one of the first steps in the silicon fabrication process. A variant of CVD commonly used in the microelectronics industry is plasmaenhanced CVD (PE-CVD), whereby a plasma of the reacting gases is created to enhance reactivity and allow faster deposition rates at lower temperatures. Several materials can be deposited by CVD or PE-CVD, most commonly polysilicon, silicon nitride, silicon oxide, and some metals. The resulting films are usually conformal, their properties depending a lot on process parameters: temperature, pressure, and gas flow rates. A CVD variant called metalorganic CVD (MOCVD) can be used for deposition of thin and thick conformal metal layers, such as in CVD metal fabrication [24]. • Atomic Layer Deposition A thin-film deposition technique related to CVD is the atomic layer deposition (ALD), whereby several precursor reactive gases are alternatively brought in contact with the substrate. Atoms adsorb on the surface of the substrate and

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Figure 1.10 Difference between wet isotropic and anisotropic etches. The under-etch present in the case of isotropic etching processes is highlighted. The slow-etching crystal planes in the case of an anisotropic etch create angled side walls, with practically no under-etch.

react with the previous layer, each step being self-limiting and leading to the growth of a new molecular layer with every cycle. By comparison to CVD, ALD is a very slow process. However, it creates fully conformal and pinhole-free films with excellent thickness control [25], which make it a key deposition technology that can enable extreme electronic device miniaturization. ALD is capable of deposition within channels and pits with large aspect ratio (including dead-end), making it very attractive for microfluidic channel coating applications. ALD deposition precursors have been developed for a large number of materials, ranging from oxides to nitrides, to metals and to semiconductors. The deposition temperatures can be maintained reasonably low (200o C–300o C), making ALD compatible with most silicon-glass microfabrication processes and even allowing the processing of some polymeric materials. • Electroplating Electroplating, or electrodeposition, is a metallization technique by which metallic ions from an electrolyte bath deposit onto a substrate, under the influence of an electric field. The substrate (or region) to be plated has to be conductive, acting as a cathode. The anode can be made out of the material to be plated, in which case it is sacrificial. Electroplating allows the deposition of thin and thick films, which have very little residual mechanical stress. It is extensively used for fabricating metallic microparts, through the LIGA process (see Section 1.2.2.3).


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1.2.5.3 Subtractive Processes

• Liftoff Liftoff is a physical process that consists of dissolving a sacrificial underlayer, typically a photoresist film, to remove material deposited on top. The liftoff process consists of casting a photoresist layer on a substrate, patterning it with a negative of the final pattern using photolithography, depositing additional layers over the patterned substrate, and placing the resulting structure in a bath of solvent for the particular photoresist. Usually energy in the form of ultrasonic vibrations is required to create cracks in the material located on top of the photoresist and allow solvent penetration. Liftoff is a versatile process, allowing the deposition of most materials on almost any flat substrate. It does not involve aggressive chemical etches, which may be incompatible with other materials already on the substrate (the typical solvents used for dissolving photoresists are well tolerated by most other materials used in microelectronics). There are, however, disadvantages of the process: all depositions have to be done at a low enough temperature to avoid crosslinking of the photoresist, whereas chips resulting from material fracturing and liftoff may stick to the substrate (sometimes permanently), compromising it. • Wet Etching Wet etches involve the immersion of the wafer in a bath of liquid etchant, which chemically reacts with the material to be removed. Good process control usually requires agitation of the etchant. Wet etchants must be chosen to have a good selectivity for the target material; well-characterized etchants for a wide variety of materials exist [26]. When applied to crystalline substrates, wet etches can be divided into isotropic and anisotropic, depending on the etch rate difference of different crystalline planes: anisotropic etches have preferred crystalline etching directions, whereas isotropic etches remove material equally in all directions. Most wet etches work very well when applied to thin films of material. However, the isotropic/anisotropic nature of the etchant has important consequences when the etch is used for bulk machining: while anisotropic etches result in structures with well-defined geometries and planar facets along predefined crystalline orientations (Figure 1.10, left), isotropic etches create structures with “round” corners and lateral dimensions larger than the original etch mask (etch bias—Figure 1.10, right). Undercutting the etch mask in bulk isotropic etches can lead to etchant trapping

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Figure 1.11 Examples of through-hole and channel profiles obtained by HFetching a Pyrex glass. The pictures show cuts through the final devices that consisted of three anodically-bonded layers: two external Pyrex wafers sandwiching a thin silicon wafer.

regardless of the amount of agitation applied, resulting in poorly controlled etching speeds. In addition, the undercut etching mask can bend or break, again with significant effects on the final etched geometry. When dealing with bulk machining of silicon, therefore, anisotropic etches are preferred: a standard etching recipe used for bulk silicon micromachining involves an aqueous solution of KOH 30% w/w at 80o C. The depth of the etched structure can be controlled either by the etch time or by using an etch-stop layer of SiO2 . For glass wet-etching the available processes are limited to the isotropic kind, a common recipe involving an HF solution in water, 50% w/w at ambient temperature. Figure 1.11 shows an example of a fluidic port and of a channel, both fabricated simultaneously by double- and single-side HF-etching of Pyrex glass (both pictures show three-layer glass-silicon-glass structures). The etch duration for the process in Figure 1.11 was approximately 45 minutes. The HF glass etchant solution is very aggressive, and it is very difficult to create etch masks that are pinhole-free and can withstand the long HF exposure required for etching deep troughs or through-wafer holes—usually multiple Cr/Au layers need to be deposited alternatively and protected with photoresist [27]. This issue, combined with the important health hazards associated with HF, make wet etching of glass a delicate process. If possible, it is wise to replace it with other glass machining techniques such as powder-blasting or ultrasonic machining, or even recently developed deep reactive ion etching glass processes. Wet etching does not require sophisticated equipment to run, and


27

thus has the advantage of being relatively inexpensive. However, in many applications involving thin films, wet etches have been replaced by dry plasma etches, which are more controlled and reliable, and are environmentally friendly by not creating massive amounts of chemical waste. • Reactive-Ion Etching Dry etching is achieved using reactive gases usually ionized and energized using a low-pressure RF plasma, hence the name reactive-ion etch (RIE). The ions in the plasma are accelerated toward the substrate by electric field created by the bias voltage on the wafer platter. The motion of the ions is mostly in a direction perpendicular to the wafer plane, leading to possible anisotropy in the etch profile. Etch recipes for a variety of materials exist, but most typical ones are for etching polysilicon (using an SF6 plasma), photoresist, and other organic materials (O2 plasma), and silicon / silicon oxide (CF4 ). A RIE variety commonly used for bulk silicon micromachining is deep reactive ion etching (DRIE). This class of processes creates very anisotropic etch profiles with nearly vertical side walls. The Bosch DRIE process is commonly used in the microfabrication industry, and it consists of two steps alternating at few seconds’ interval: 1. A nearly-isotropic silicon RIE step (e.g., using SF6 gas); 2. A passivation step (e.g., using C4 F8 ), which deposits a Teflonlike material on the walls. Each RIE step first removes, by sputtering, the passivation layer at the bottom of the etched structure, then reactively etches a few micrometers of silicon. The lateral passivation layer is removed more slowly, assuring a nearly vertical etch at each step. Figure 1.12 shows an example of a microfluidic channel and through-hole fabricated by successive front- and backside DRIE. Depending on the relative durations of the passivation and RIE steps, walls that deviate from the vertical can also be achieved. Some scalloping of the walls may result as a consequence of the alternating etch-passivation nature of this process. DRIE processes for etching glass (Pyrex) substrates, albeit very slowly, have been reported [28]. RIE processes can be combined with grayscale lithography or photoresist reflow to create custom three-dimensional shapes in silicon in a single processing step. The exposure characteristics of the resist used, as well as the selectivity of the etch, need to be very well known to assure accurate shape rendering in grayscale processes.

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Figure 1.12 Example of a channel and through-hole fabricated on a silicon substrate by DRIE of the front, and respectively, the back side of a silicon wafer.

Alternatively, certain particularities of plasma etch processes can be used to create smooth three-dimensional shapes: since the etch depth depends on the size and density of the apertures being etched, a surface with smoothly varying topography can be realized by creating a template with apertures arranged in specific patterns, then performing the dry etch, and finally performing an isotropic etch after removing the template to smooth the surface. This process is called MEMSNAS [29], an acronym for microloading effect for micromachining 3-D structures with nearly arbitrary shapes. It can be used to create complex geometries, particularly in situations where a large curvature radius is desired, such as in fabricating microlenses (Figure 1.13) or electrostatic actuators. Alternatives using DRIE anisotropic etching have been developed as well, in which case the depth of the resulting structures depends on the size of the apertures in the template [30]. • Powder Blasting Powder blasting (or sandblasting) is a process that has been used for many years in conventional manufacturing to texture and deburr machined parts. It involves a jet of small particles (Al2 O3 is a common powder material), carried by a high-speed gas stream, which impact and erode the part. The process is particularly effective on brittle materials such as glass, silicon, and ceramics, where the impact with powder particles creates microcraters, slowly removing material. Typically a mask is used to confine the powder


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(c) Figure 1.13 MEMSNAS process showing schematically the desired final shape (a), an aperture template (b) that can be used to generate the shape, and a similar structure realized using MEMSNAS (c). (Reproduced with permission from [29].)

stream to well-defined substrate regions. The mask can be deposited on the substrate wafer photolithographically, prior to the blasting process, or may be a metallic shadow-mask maintained in mechanical contact with the substrate. Modern powder-blasting equipment can utilize powders with very fine particles (down to a few micrometers), which allows the fabrication of well-defined features with lateral dimensions as small as 10 micrometers. Powder-blasted surfaces typically have a roughness on the micron scale, rendering this process unsuitable for certain optical applications without subsequent treatment. The lateral walls have slightly slanted walls, with a taper typically of about 5 to 15 degrees, which limits the aspect ratio of powder-blasted structures to below 10. Multistep and double-sided blasting is possible, allowing complex three-dimensional geometries to be manufactured. This technique is particularly interesting for manufacturing cavities and through-holes in glass wafers (Figure

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Figure 1.14 Cavities of different diameters machined in parallel by powder blasting a glass wafer. (Used with permission from Anteryon B.V., www.anteryon.com.)

1.14). • Ultrasonic Machining Ultrasonic machining is a process similar to sandblasting, in the sense that it works by generating microfractures in brittle materials like glass, silicon, and ceramics. However, instead of using powder granules to achieve this goal, ultrasonic machining utilizes a metallic tool vibrating at frequencies of tens of kilohertz, and in close proximity to the substrate being machined. An abrasive slurry is applied to the part, serving the double purpose of speeding up the machining, and carrying the removed material away from the part. Modern ultrasonic mills can produce very small features, down to a few tens of micrometers, while keeping the substrates scratchfree for subsequent microfabrication steps. Aspect ratios of 1:5 can be achieved for small through-wafer holes, and wafers as thin as a few tens of micrometers can be processed by this technique. • Polishing and Planarization Polishing has long been regarded as a dirty process, involving particulates and contaminants, and thus unsuitable for the microelectronics industry. In more recent years, however, chemical-mechanical polishing (CMP) has been adopted as a standard manufacturing step in many microelectronics and MEMS fabrication processes. The process consists of placing the wafers against a spinning polishing pad covered with a layer of silica slurry. The chemical composition and the pH of the slurry are controlled. The role of the chemicals in the slurry is to react with the surface


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of the substrate and make it susceptible to mechanical abrasion; chemical composition needs to be altered depending on the nature of the substrate. CMP is capable to planarize wafers of small as well as large diameters, and can produce low surface roughness down to a few nanometers. In microfluidic applications, CMP can be used to reduce the thickness of a silicon wafer after it has been bonded to a different substrate. Devices integrating very thin mechanical silicon structures can be achieved with this method, which would have been impossible to process otherwise. For MEMS and microfluidics, CMP can also be used in a manner more closely related to its microelectronic applications: to planarize dielectric layers deposited over metallic traces for interconnects [31], particularly in view of subsequent bonding processes that may require high planarity. Other means of planarization include depositing a glass layer, either by spin-coating or by sputter deposition, which is then reflown to create a locally planar structure. The degree of planarization achievable by the reflow technique is poorer than that of CMP. • Focused Ion Beam Machining Focused ion beam (FIB) etching can be used to remove almost any materials. The highly energetic ions are produced by field emission from a tip, and are focused down to a beam only a few nanometers in diameter. The beam can be steered to machine away, and sometimes to deposit, material onto a substrate. The resolution of the patterns that can be machined using FIB is on the order of a couple of nanometers; however, the process is very slow and expensive. It is therefore used mostly to repair small areas of semiconductor wafers, to perform punctual analyses on certain parts of a device, or to patch lithography masks. 1.2.5.4 Formative Processes

• Annealing Annealing is a thermal process used to alter the state of a material. In an annealing step, the substrate is heated on a hotplate or in an oven, maintained at a suitable temperature, and then cooled. During the annealing process atoms have higher mobility, and can rearrange in more stable configurations. Therefore, annealing a thin film helps to reduce the stresses that may have appeared during deposition. The microstructure of the film can also be affected—significant grain growth during annealing can be observed for some materials, like sputtered or evaporated gold films, affecting their mechanical

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and electrical properties. Sometimes annealing is performed in a controlled atmosphere, such as in semiconductor diffusion doping or oxidation. Photoresist spincoating and lithography processing routinely use annealing steps to cure and/or crosslink the polymer layers, or to cause photoresist reflow, a process used in threedimensional fabrication of optical microcomponents. Reflow annealing steps are often used for reflowing glass or dielectric layers used in microelectronics as a first step in device planarization. Device fabrication processes are usually followed by an annealing (temperature aging) step, usually at a temperature significantly higher than the maximum working temperature of the finished product. This assures stability of the device under operation conditions, minimizing drifts that may otherwise develop over time in the device being fabricated. • Doping Doping is a widely used process in microelectronics and microfabrication. It involves the enrichment of the silicon substrate with dopant atoms having a different number of valence electrons, thus either donating an electron (n-type dopants such as phosphorus or arsenic) or creating a “hole” (p-type dopants such as boron). The doped silicon thus becomes more conductive. Doping is used routinely in microelectronics to create the active devices used in integrated circuits. It is also used for microfabrication, particularly to create doped polysilicon elements (used in piezoresistive strain gauges), or as an etch stop in wet etching. Two main processes are used for doping silicon: diffusion and ion implantation. The former is achieved by annealing at high temperatures in a controlled atmosphere, being relatively inexpensive, and the latter can be performed at room temperature and is more controlled; however, it requires the use of expensive facilities and equipment. It is important to note that the level of impurity atoms, and hence the conductivity of the silicon substrate, depend not only on the doping level but also on the fabrication process used for the silicon ingot. The least contamination is obtained using the float-zone fabrication method; the more popular Czochralski method introduces more contamination and results in wafers with lower resistivity. • Oxidation Oxidation of silicon is a critical step in microelectronics fabrication, as it is the most common dielectric layer used in integrated circuits. Different methods of producing silicon oxide on a wafer exist: thermal wet and dry oxidation as well as local silicon oxidation (LOCOS), and CVD. In dry oxidation, the wafers


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are annealed in a dry oxygen atmosphere. This leads to a highquality oxide; however, the diffusion process is very slow and leads to very long processing times, particularly if thicker oxides are desired. Because of this inconvenience, wet oxidation (where water vapors are added to the atmosphere) is often preferred when growing thick oxides. An initial and final thin dry oxide may be grown as well, to improve the dielectric quality of the oxide layer. Both types of thermal oxidation are typically performed in quartz-lined furnaces, at temperatures close to 1,000o C. Because of the high temperatures involved, thermal oxidation is best performed early in the fabrication cycle to avoid incompatibility with previous fabrication steps. CVDdeposited oxide is usually of lower quality, contains contaminants, and does not have a very good dielectric strength; it can, however, be deposited at a lower temperature than thermal oxide. During plasma etching processes, especially when a high bias voltage is used, the oxide layer can undergo dielectric breakdown, so care must be taken to avoid such treatments. Often, the LOCOS process may prove useful to pattern the growth of thermal silicon oxide; it is achieved by masking the rest of the wafer with a silicon nitride layer, which acts as a diffusion barrier for oxygen. 1.2.5.5 Bonding Processes

• Anodic Bonding Anodic bonding is the oldest method developed for microfabrication applications, and remains the method of choice for permanently bonding sodium-rich glasses to silicon substrates. The surfaces to be bonded are placed in a vacuum chamber, brought into contact, and then heated to a temperature of 300o C to 500o C. An electric field is created by applying a voltage of several hundred volts between the two substrates (silicon being the cathode), which causes the highly mobile sodium ions to diffuse into the silicon. The resulting electrostatic attraction initiates the bond, which is enhanced by the formation of permanent chemical bonds between the silicon and the glass. The bonding current is normally monitored—complete bonding is achieved when the current vanishes. The bonding can also be monitored optically by watching the color change of the silicon wafer—because of the optical index-matching of Pyrex and silicon oxide, once the wafers are bonded, the color specific to the silicon oxide thickness preexisting on the wafer disappears. Since bonding is performed at a relatively high temperature, care needs to be taken to use substrates with matching thermal expansion coefficients.

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Several types of glass qualify—Borofloat manufactured by Schott and Pyrex 7740 manufactured by Corning are common choices. While bonding can be achieved reliably through buried oxide layers such as those present in SOI wafers, oxides present on the surface of silicon may affect the bond and therefore need to be kept thin (below approximately 200 nm). Figure 1.11 shows an example of a threelayer structure bonded by a two-step anodic bonding process. While using a transparent substrate such as a Pyrex wafer has advantages in many applications (most notably because of optical transparence for visible light), it is not a requirement for anodic bonding—two silicon wafers can also be anodically bonded by using a sodium-rich glass interlayer, deposited in thin film by spin-coating, sputtering, or evaporation. The anodic bonding method can also be used for joining other materials, most notably glass to metals. • Eutectic Bonding Certain binary alloys can exhibit melting temperatures that are significantly lower than those of either component in pure form. The alloy composition at which the minimum melting temperature is reached is called the eutectic composition of the mixture, and the corresponding temperature is called the eutectic temperature. This property is used in the eutectic bonding of silicon wafers using a gold thin-film interlayer. Silicon forms an eutectic alloy with gold (97% Au, 3% Si), with a melting temperature of 363o C. As the temperature of the compressed Si-Au-Si stack is raised past the eutectic point, diffusion of the gold atoms into the silicon substrate is enhanced, leading to the eutectic alloy formation. As the temperature cools, the eutectic alloy solidifies, serving as a solid bonding layer between the silicon wafers. While fairly common in the MEMS world, eutectic bonding has lower yield than anodic bonding. • Direct Bonding Techniques Bonding between similar materials without using an intermediate layer can be achieved by direct (or “diffusion”) bonding. The method involves extensive cleaning of the substrates to remove all surface contamination (organics, particulates, oxide layers), bringing the substrates in intimate contact, and annealing at very high temperatures for a number of hours. When applied to silicon, the method is also called “fusion bonding.” Due to the elevated processing temperatures (approximately 600o C for glassglass and 1,000o C for silicon-silicon bonding), these bonding techniques often cannot be used for the packaging of complex, integrated devices (especially those involving low melting-point metals such as aluminum in their fabrication), which limits their utility in many


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applications. Multilayer glass microchips with integrated metallic electrodes can, however, be manufactured using this technique. Direct bonding (particularly for joining silicon substrates) is very intolerant to nonplanarity of the wafer due either to roughness or to particulate contamination, and thus requires ultraclean facilities. Bonding defects in silicon-silicon structures can be difficult to locate optically, and usually the process requires postbonding inspection using infrared microscopy equipment. • Adhesive Bonding Adhesive bonding is one of the most versatile techniques used in macroscopic packaging, due to easy processing and to its ability to bond a wide variety of materials. It can also be applied in microfabrication, particularly through the use of photopatternable adhesives. These are typically deposited in thin films onto a surface, patterned using standard photolithographic techniques, followed by the mechanical joining of the substrates to be bonded. The two most popular adhesives for this purpose are photosensitive benzocyclobutene (BCB) and SU-8. Alternatively, optical adhesives can be allowed to wick between the wafers to be bonded by capillarity, and subsequently UV-cured [32].

1.3 Technology Selection Criteria As it has become apparent from the previous section, a large array of technologies are available to the designer of microfluidic and/or MEMS devices. The choice of technology that is appropriate to a specific application is not trivial. Several criteria need to be evaluated, the choice usually involving a compromise specific to the application at hand. Normally, the highest level of integration is desired in the final device—one would like to have a fully contained device that is small, functional, cheap, and compatible with the end user’s environment. Sometimes achieving this goal is not easy, or may even be impossible—in such cases hybrid approaches need to be developed, involving multiple interconnected devices, and manufactured using different technologies. When it comes to making a technology choice, there are no rules cast in stone. This section is only intended as a rough guide, providing a somewhat systematic approach to what could otherwise seem like a technology jungle. An introduction to a number of technology solutions chosen in commercial applications is provided in Chapter 5.

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1.3.1 Required Functionality The first criterion to be analyzed is that of the functionality required by the application at hand. A functional diagram of the device helps determine the building blocks required. Often, several diagrams can be devised for the same application utilizing different functional blocks. Since some functional diagrams may apply better to certain manufacturing technologies than to others, it is a good idea to identify many potential functional implementations for any given device. This brainstorming phase assures a lot of flexibility in the design phase, and will reflect in the quality and cost of the final device. Figure 1.15 provides an example of different functional diagrams pertaining to the same final application: the detection of a specific property of a liquid sample extracted from an external liquid stream by means of a chemical reaction with a specific reagent. The detection is of optical nature— we assume the chemical reaction leads to a change in the absorbance of the sample at a particular wavelength, and that the change in absorbance is related to the particular property we want to measure. This example may pertain to an environmental monitoring application (e.g., the water quality in a river, in the vicinity of a plant, needs to be continuously monitored for presence of an industrial contaminant). Several devices need to be installed permanently upstream of the plant to get a baseline, and several downstream to measure the contamination from the plant relative to the baseline. The top diagram in Figure 1.15 involves two liquid sources: a sample inlet and a reagent reservoir. The sample is filtered, and both liquids are then pumped, using independent micropumps that are regulated to maintain a given reagent concentration. The combined streams are then mixed using a micromixer, pass through the detection region, and then the fluids are discarded. In this application, the requirement to have two micropumps results in increased cost and complexity, and it also requires the pumps to be placed upstream of the detector region. This may not be desired, since the pumps may have large internal dead volumes that may dramatically influence the response time of the sensor. Since many devices need to be produced at low cost, the ideal fabrication technology may involve clear plastic fabrication (by injection molding) using integrated piezoelectric micropumps. Piezoelectric micropumps require very high voltages, however, which may pose a problem if the sensors are going to be permanently installed on the riverbed. Piezoelectric micropumps are also relatively power-hungry, and may require frequent battery changes. The middle diagram in Figure 1.15 represents an alternative involving a single micropump (and thus reducing both cost and power consumption relative to the previous design). In this case, the concentration of reagent in the mixed stream is regulated passively, by using flow restrictors. In this


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Figure 1.15 Three possible functional diagrams for the same final water quality monitoring application.

configuration the micropump can be placed downstream from the detector, eliminating the pump dead volume issues encountered in the previous design. This is a good option as long as the flow resistance of the filter remains constant over the lifetime of the sensor. If there is significant filter contamination, however, the concentration of reagent may drift in time, leading to a compromised measurement. The filter contamination may be eliminated by making sure that only a very small portion of the liquid stream is sampled and the rest exhibits a cleaning action on the filter (a configuration called tangential, or cross-flow filtration). Again, clear plastic fabrication may be the ideal way to go in this case. Finally, the bottom diagram in Figure 1.15 involves a fully passive system—no micropumps are used, flow generation being assured by a lowpressure reservoir (which may be as simple in concept as an evacuated volume connected to the outlet of the system). A normally closed microvalve controls the operation of the device, allowing flow only when a measurement is required. The concentration of reagent is also regulated passively by using flow restrictors, and the microvalve is placed downstream of the detector, all dead volume issues being eliminated altogether. Filter contamination may become an issue, again, and a self-cleaning tangential filtration design may be required. Using an external reservoir is convenient, as no reagent

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is released back into the environment after the analysis (which may raise environmental concerns for the previous designs). The total volume that can be sampled over the lifetime of the sensor is directly related to the volume of the low-pressure reservoir. After the reservoir is full, it needs to be reevacuated—the time scale corresponding to this can be of many years in case the measurement frequency is reasonably low. This design may be manufactured from any optical material (plastic, glass, glass-silicon-glass multilayer); however, fabrication in glass/silicon allows the integration of an electrostatic microvalve onto the chip, which may be preferable, from a cost and size perspective, to using an external solenoid valve. Due to the batch fabrication and high miniaturization capability of silicon/glass fabrication, the cost of an integrated silicon/glass device may be significantly lower than that of using an injection-molded plastic device fitted with either micropumps or external microvalves. Overall, the system designer is likely to select the bottom diagram in Figure 1.15 due to environmental reasons and simplicity, and choose glass/silicon/glass fabrication as the fabrication technology due to the chemical stability of the materials and the high degree of integration (and thus lower cost) possible. This choice would have not been as obvious, however, had different alternatives not been explored. 1.3.1.1

Functional Blocks

A number of essential functional building blocks are required in most microfluidic devices. These include sample acquisition and manipulation blocks (ways to control, generate, and measure flow), physical and chemical manipulation blocks (thermal treatments, mechanical manipulators, reagent injection, and mixing), and sensing and measurement blocks (stress, pressure, temperature, fluorescence, optics, electrochemistry, and so on). This section discusses the manufacturing feasibility of a few functional block examples using various technologies presented earlier in this chapter. Chapter 2 will then provide a much more detailed review of microfluidic building blocks. • Flow Control Internal flow control can be achieved by using one of several possible types of microvalves and micropumps. Microvalves can be either passive or active. Passive microvalves are used mostly in flow regulation and find important applications in the internal construction of micropumps. The majority of reported active microvalves and micropumps are fabricated using silicon and/or glass micromachining, and are


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actuated using pneumatic, thermopneumatic, electrostatic, electromagnetic, or piezoelectric means [33]. There are academic reports as well as commercial examples of micropumps fabricated out of injection-molded plastics; pneumatically controlled valves and pumps operating in peristaltic mode can be made entirely of PDMS rubber using soft lithography [34]. Metallic fabrication is, in principle, possible, using piezoelectric, electromagnetic, or pneumatic actuation. • Flow Sensing Different principles of flow sensing exist. Microfluidic implementations involving either differential pressure or thermal calorimetric and anemometric principles are being manufactured almost exclusively using silicon micromachining. Parylene-film thermal sensors using a hybrid polymer-silicon manufacturing approach have also been reported [35]. Thermal time-of-flight microfluidic sensors have been reported using silicon microfabrication, leading to a fully integrated sensor [36], or using a microchannel fabricated in any optical material combined with external laser beams for both heating and detection [37]. Particle imaging velocimetry and particle tracking can both be used to measure flow profiles and velocities in any optical material; however, they both require intrusive action to introduce colloidal particles in the flow and large external optical and postprocessing equipment. Different drag-force sensor designs have also been implemented in microfluidic systems, utilizing silicon components either as cantilevers, membranes, or beams. • Thermal Treatments and Temperature Sensing Many microfluidic applications require thermal treatments of fluid samples. In some applications heating the sample is sufficient, which can be accomplished using integrated heaters. The technology that offers the highest versatility for creating built-in heaters (with optional on-chip temperature sensing for feedback control) is silicon/glass fabrication. Any type of heater or temperature sensor compatible with microfabrication can be utilized: thin-film metallic heaters, polysilicon heaters, platinum RTDs, or micropyles, suspended on thermally isolated membranes or on silicon structures. Plastic fabrication allows the integration of heaters if the metal deposition process conditions (particularly the temperature) are compatible with the type of plastic being used; usually this restriction limits deposition to sputtering processes. As it may be obvious, builtin heaters in a plastic device are limited to operation below the melting temperature of the plastic substrate. Plastic and glass can be patterned

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with optically transparent electrodes (made, for example, of indiumtin oxide). If very localized heating is required, a thermally isolated structure or a low thermal conductivity substrate is preferred—as a rule of thumb, plastic substrates conduct heat least efficiently, followed by glass, ceramics, silicon, and metals. In addition to internal heaters, external temperature control is also possible in microfluidic devices. Due to the relatively low thermal mass of a microchip, its temperature can be cycled rapidly. Very popular devices for this purpose are thermoelectric devices (Peltiertype), which also offer the cooling possibility. When using external coolers, however, a relatively massive thermal sink is required, often significantly increasing the size of the device. In recent years, promising research on the development of highly efficient thermoelectric coolers [38] and on MEMS-like fabrication processes for thermoelectric coolers [39] has opened possibilities for even higher integration using very localized thermal control in microfluidic chips. • Mechanical Manipulators Micromechanical manipulators have been almost exclusively manufactured out of silicon since their very inception [40], using electrostatic comb drives for actuation [41]. They have many potential application in microfluidics; however, electrolysis effects due to the high potential differences between the comb electrodes make them of limited use in a liquid like water. New conformal dielectric layer deposition techniques like atomic layer deposition have a significant potential to enable this technology in the field of microfluidics. Micromanipulators can also be manufactured out of metal [42]; however, the issues of electrical routing and encapsulation in this case are more difficult. • Pressure Sensing The majority of miniaturized pressure sensors are based on a flexible diaphragm manufactured using bulk microfabrication of silicon and using either piezoresistive or capacitive sensing. Bulk microfabrication results in single-crystal, unstressed, silicon membranes, which can lead to good accuracy and stability. Even higher degrees of miniaturization and complexity can be achieved using surface micromachining [43]. There are also reports of completely passive polymer-based miniaturized pressure sensors, made, for example, of parylene [44].


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Figure 1.16 An example of plastic fluidic interconnect achieved by using a standardized port interface and closely fitting plastic parts. (Used with permission from [45], www.thinxxs.de.)

1.3.1.2 Interconnects

The subject of interconnects is central to almost any microfluidic application. Two types of interconnects are critical to most microsystems: fluidic and electric. Fluidic interconnects are required to transfer liquid or gas samples from one device to another, whereas electrical interconnects serve the function of transmitting signals and power between the microfluidic system and the outside world. • Fluidic Interconnects The most important functions that need to be addressed by fluidic interconnects are to minimize dead volumes and to be able to seal reliably throughout the working temperature and pressure range of the device. For soft elastomeric devices, sealing is most commonly achieved by pushing a capillary tube (typical outer diameter 1 mm) through a hole premade in the elastomer. By making the hole in the elastomer of a slightly lower diameter than the diameter of the tubing, one can achieve a reasonable sealing up to pressures of a few bar. Alternatively, plastic connection ports can be adhesively attached to PDMS, usually after another surface activation step. Metallic microfluidic devices can be manufactured to accept several different fluidic connection options: O-rings, swaged conical ferrule fittings, adhesively coupled fittings, small OD capillaries adhesively attached inside edge-terminated microchannels, brazed tubes, and so forth. Ferrule-type fittings, swaged on 1/32 inch or 1/16 inch diameter metallic capillary tubing, as well as metallic tube brazing, offer the

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Figure 1.17 An example of commercial plastic fluidic ports (manufactured by Upchurch Scientific) adhesively attached to a silicon-glass microfluidic chip.

best pressure and temperature capability (in excess of 1,000 bar, 200o C), as well as reasonably low dead volumes. Plastic devices are usually connected using adhesive techniques, by using close-fit matching plastic parts (Figure 1.16 provides an example), or by connecting soft capillary rubber tubing around protruded injection ports that have been injection-molded or laserwelded. Silicon and glass devices offer a reasonable array of interconnection possibilities. Commonly, commercial plastic fittings are adhesively attached to ports located on the surface of the chips (Figure 1.17). This is usually a good option for larger chips, and can usually withstand several tens of bar; at higher pressures, the adhesive connections often fail. Space constraints eliminate this possibility for chips that are smaller than approximately 1 cm square. Microfluidic chips may be attached to other chips or to fluidic manifolds using face-seal O-ring connections. Properly engineered, these can assure reliable operation up to several hundred bar. Alternatively, by metalizing the area surrounding a fluidic port, solder-based connections capable of withstanding pressures of a couple hundred bar can be made either to metallic tubes or to other chips [46]. Kovar metallic tubes can also be bonded to glass [47], a technology long used for military applications. Finally, thin glass capillaries can be adhesively bonded within edge microchannels, creating a very low-dead-volume connection. • Electric Interconnects Electric interconnects in microsystems are as important as fluidic ones, with many active elements within a microsystem being electrically driven. Thin metal films cannot be reliably deposited on elastomers, so chips manufactured by soft lithography need to rely on electrodes


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deposited on a different substrate. Typically, electrodes are made on a glass slide, and the PDMS microfluidic chip is then aligned with the electrical structures and bonded to the glass [48]. Two disadvantages of this approach are that the electrodes are directly exposed to the fluid and that PDMS does not usually bond to the metallic electrode, potentially leading to the development of leaks. To overcome both of these inconveniences, a thin PDMS film can be cast on the glass slide, and the remaining PDMS structures can be bonded to this film. Metallic chips are normally conductive, so electrical interconnects cannot be incorporated unless dielectric treatments are applied to the metal, which will affect its capability to bond (via brazing, regular welding, or solid-state welding) to another metal part. Silicon and glass fabrication can borrow many types of interconnect technologies from the microelectronic industry. There are aspects of the interconnects that are specific to microfluidic applications and stem from the requirement of encapsulation. In particular, wafers need a high degree of planarity to facilitate bonding, so techniques such as chemical-mechanical polishing CMP planarization [31] and glass reflow planarization [49] need to be used for the purpose of creating hermetic interconnects. It is interesting to note that in microfluidic applications the hermeticity requirements stem from a need to contain the fluids within the microsystem, whereas in microelectronics the requirement is often reversed, hermetic interconnects being needed to isolate the electronics from potentially aggressive external environments (e.g., in permanent implantable installation [50]). Adhesive bonding can be achieved regardless of planarity and has been used in the manufacturing of integrated siliconglass systems [32].

1.3.1.3 Integration

When attempting to create an autonomous system that integrates multiple sample preparation and analysis functions, the system architect is faced with two options: one is to create a system where individually manufactured components are interconnected, resulting in a completely functional system (hybrid integration); the other is to choose a single fabrication technology that allows the monolithic integration of all desired functions in a single device (monolithic, or device-level integration). Trade-offs exist for either choice, and the decision can be different from one application to the next. Several examples of commercial approaches to integration are given in Chapter 5.

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From a device-level integration perspective, two technologies stand out among the different options discussed in this section: silicon/glass microfabrication and multilayer soft lithography. The versatility of silicon as a material, capable to act simultaneously as a substrate for microelectronics and as a structural mechanical material, make it the material of choice in a large number of applications. A greater number of reliable well-established processes exist for manufacturing silicon and glass than for any of the other material considered in this section. The interconnect technologies developed in the IC industry for routing signals between different areas of a chip can also be applied to microfluidic applications, leading to reliable processes for integration and encapsulation. As the choice material in microelectronics, the silicon substrate can also incorporate analog signalconditioning electronics or even a digital electronics front end. The ultimate degree of miniaturization and integration of silicon-glass microfluidic systems will not be determined by the limitations of fabrication processes, but rather by physical parameters such as the minimum size required to perform the intended function (for example, a system for manipulating cells of a certain type must be designed with flow channels at least as large as the cell in question), the size of the fluidic connections, and the requirement to have bonding pads for electrical lines. High levels of fluidic control integration can equally be achieved with multilayer soft lithography. The soft and elastic nature of the cured PDMS material, as well as its optical transparency in the visible and UV, make it an excellent choice in a range of low-pressure applications. The possibility to manufacture soft channels that overlap and are separated by thin flexible membranes allows the implementation of a number of pneumatically (or hydraulically) controlled fluidic functions such as on/off flow control and peristaltic pumping. What is more interesting is that the pneumatic or hydraulic control of multiple networks of channels can be multiplexed [51], leading to complex flow manipulation and control capabilities (Figure 1.18). Devices integrating multiple electrically controlled functions have been manufactured using hard polymeric materials as well, parylene being a material of choice for its compatibility with clean room fabrication processes [35]. The level of integration complexity that has been achieved with parylene fabrication, while very promising in some application areas, is currently lagging behind that of silicon or PDMS.


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Figure 1.18 An example of high-level fluidic control integration using soft lithography in PDMS and control line multiplexing. (Reproduced with permission from [51].)

1.3.2 Physical Requirements 1.3.2.1 Optical Requirements

Chemical or biological processes taking place in a microchannel are often monitored using optical means. From simple visualization, to fluorescence detection, to through-chip spectroscopy, and to single fiber spectroscopy, optics is extensively used in conjunction with microfluidics. The most common optical materials used in microfabrication are glass, glass-siliconglass hybrid, injection molded or embossed transparent plastics such as PC and COC, and PDMS. While not optically transparent in the visible range, high resistivity silicon (such as that grown using the floating-zone technique) is a very good material for infrared transmission. For through-chip optical visualization, visible-UV, or fluorescence spectroscopy, channels etched fully through the thickness of a silicon wafer can also be encapsulated using two bonded Pyrex covers.

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1.3.2.2


Wetting Requirements

The wetting properties of various materials are very important in microfluidics, particularly when dealing with two-phase flows such as in digital microfluidics. Microfabrication-compatible materials that are hydrophobic (and oil-wet) include PDMS, many hard plastics, and silicon. Silicon, however, quickly develops a thin native oxide layer in contact with oxygen, which changes its wetting properties. Glass, being a polar material, is hydrophilic, and so is oxidized silicon. Metals usually have mixed wetting properties. Many physical and chemical ways exist that can be used either to uniformly modify, or pattern the wetting properties of different materials. 1.3.2.3

Temperature

Microfluidic systems are often required to operate at elevated temperatures. The manufacturing technique adopted will be the main factor dictating the maximum operating temperature of the device. Plastics, as a rule, have low temperature resistance (typical heat resistance of plastics commonly used in injection molding ranges from approximately 100o C for polystyrene to about 140o C for polycarbonate). There are, however, high-performance plastics such as PEEK that can hold mechanical strength to temperatures in excess of 150o C, and their use in microfluidic applications is starting to be investigated [17]. PDMS is a cross-linked elastomer, and consequently can withstand higher temperatures than injection-molded plastics. Operation up to 200o C does not pose significant problems. Silicon / glass microchips that do not embed electronic components typically allow operation up to the temperature of the last high-temperature step used in the microfabrication process. For a typical anodically-bonded microchip, this corresponds to approximately 300o C–450o C. These temperatures, again, are higher than typical requirements of microfluidic applications. The situation changes if the microfluidic chip contains microelectronic circuits as well—in this case, the maximum operating temperature is limited by the specific electronic circuit, and can be as low as 85o C (the typical rating of consumer electronic components). In case the design includes thin metal film heaters (such as in a chip that has the function to generate periodic bubbles in a liquid, an arrangement commonly encountered in bubble-jet printing heads), the local temperatures at the heater may easily reach the melting point of the metal. Heater temperature regulation is required in such cases. Very high operation temperatures can be achieved with metallic microchips, where the limitation is usually given by the melting temperature


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of the metal, which can be in excess of 1,500o C for a material like titanium, and thus completely covering all possible requirements of microfluidic applications. 1.3.2.4 Pressure

Due to their intrinsic “soft” nature, elastomeric materials such as PDMS cannot be used in applications requiring high pressures; as pressure is increased beyond a few bar, the cross-section of the channels changes from rectangular to round, bonded layers tend to delaminate, and the tubing used for attaching fluidic connections tends to pop out of the device. Plastic devices (nonelastomeric) show a better performance; however, their utilization for high pressure applications is not recommended either. The best performance in pressure is achieved with metallic microchips. Diffusion bonded metallic parts with microchannels embedded at the interface can easily withstand hundreds and even thousands of bar. Depending on the bonding technique used in manufacturing and on the fluidic interconnect solution adopted, silicon/glass fabrication may allow operation up to a few hundred bar. At such pressures, the brittle nature of the materials starts playing a key role: crack propagation becomes the dominant failure mode. Usually, however, the bonding interface delaminates or the fluidic connections to the devices break before catastrophic device failure. 1.3.3 Chemical Resistance Many microfluidic applications deal with relatively mild liquids such as water, in which case almost any material described in this chapter can be chosen for fabrication. There are, however, situations where more aggressive fluids need to be used: acid or alkaline solutions, different types of oils, solvents, and so on; in such cases the proper choice of materials can become critical. A chemical compatibility chart always needs to be consulted before a manufacturing process or material is selected for a specific application. PDMS is a relatively inert material from a chemical perspective. However, it is very permeable to most organic solvents (which also swell it), to most gases, and even to water, which precludes its use in certain applications. While normally considered a nuisance, the permeability of PDMS can be put to good use in certain applications [52]. Most plastics suitable for microfabrication are usually well suited for working with acids; however, they show poor resistance to most organic solvents. Notable exceptions are the thermosets, which usually swell but do not dissolve when used with solvents. Exceptions are also some high-performance plastics, such as PEEK,

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and fluorinated compounds like Teflon, which exhibit outstanding chemical stability. Metals, as well as silicon and glass, are usually the best choices when dealing with organic solvents. However, they are of limited utility in acid environments, which are known to etch metals and glasses. Similarly, alkaline solutions, particularly KOH, are etchants for silicon as well as for other materials commonly used in microfabrication. Conformal coating technologies such as atomic layer deposition show high promise in improving chemical resistance of microfluidic devices. Since an integrated microsystem fabricated using silicon/glass technology typically involves multiple materials that will be in contact with the fluid, consultation of an etch chart for microfabrication processes [26], and of chemical compatibility tables, is highly recommended prior to embarking on an application-specific design. 1.3.4 Cost of Production Cost is an important factor that affects the market viability of any product in any domain. Lower fabrication costs can sometimes enable new market channels, often also leading to new choices of technology. For example, a specific device may be fabricated using an expensive technology that has proven long-term reliability, but may also be manufactured using a more economical process to create a single-use disposable product. Many factors play a role in the cost of a microfluidic system: materials, manufacturing, packaging and testing costs, and so forth. PDMS manufacturing, while interesting for lab usage and prototyping due to its simplicity and low equipment cost, requires a large amount of manual intervention and therefore is not an ideal commercial alternative. Additionally, the relatively large size of PDMS chips (typically on the order of centimeters) makes batch fabrication of PDMS devices problematic. Plastic manufacturing, either by hot embossing or by injection molding, is among the least expensive technologies for manufacturing microfluidic parts in large quantities. Tens of thousands of parts can be manufactured daily in a fully automated process. It does have high upfront costs, however, in the manufacturing of the mold inserts. Some applications can be tackled using less standard fabrication techniques such as paper manufacturing. This technology has the potential to bring the cost of certain microfluidic applications down by several orders of magnitude, particularly for single-use applications in the medical and environmental analysis industries. Metallic fabrication using batch processes such as LIGA can create microscale parts very economically. However, for larger components such as


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a full microfluidic system, it starts to lose its appeal and quickly becomes overwhelmingly expensive. An interesting alternative that also offers batchfabrication capabilities is the bonding of photoetched metallic stacks—this technology allows metallic parts to be manufactured relatively economically in large quantities, and is used in the fabrication of microreactors. Other, more traditional machining techniques for metal are serial in nature, and thus result in very high cost per part. Silicon and glass microfabrication allows batch production of devices; however, the cost of a full manufacturing run can be intimidating. Unit cost depends on the degree of miniaturization achieved and on the number of wafers processed in parallel. The cost per part may drop as low as that of injection-molded components, if a reasonably small device footprint, on the order of few millimeters on the side, can be achieved in production. Given the potential of this technology to monolithically integrate multiple functionalities, it can further reduce cost by eliminating external components.

1.4 Conclusions We have reviewed the most popular techniques for manufacturing microfluidic components. We looked at different substrate materials, and at several techniques for machining each type of substrate. The reader should now have a solid understanding of the most common fabrication processes and of their relative advantages and drawbacks. The technologies that have achieved the highest degree of devicelevel functional integration so far are silicon/glass microtechnology and multilayer soft lithography. Hybrid integration approaches, on the other hand, incorporate several functional blocks fabricated using a wide array of different technologies such as elastomer, plastic or metallic fabrication, and silicon micromachining. The selection of the right technological approach for a specific application is a complicated process, and sometimes it can be subjective—we tend to make the choices we are most familiar with, even though they may not be optimal for the problem at hand. Therefore, we tried to provide some objective guidelines for technology selection by concentrating on what we see as the most important criteria: required functionality, environmental compatibility (from both a mechanical and a chemical point of view) and cost.

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References [1] Plummer, J., Deal, M., and Griffin, P., Silicon VLSI Technology: Fundamentals, Practice and Modeling, Upper Saddle River, NJ: Prentice Hall, 2000. [2] Nguyen, N., and Wereley, S., Fundamentals and Applications of Microfluidics, Norwood MA: Artech House, 2002. [3] Hsu, T., MEMS & Microsystems: Design and Manufacture, Boston, MA,: McGraw-Hill, 2002. [4] Maluf, N., and Williams, K., Introduction to Microelectromechanical Systems Engineering, Norwood MA: Artech House, 2004. [5] Groover, M., Fundamentals of Modern Manufacturing: Materials Processes, and Systems, New York, NY: John Wiley and Sons, 2007. [6] Beeby, S., Ensell, G., Kraft, M., and White, N., MEMS Mechanical Sensors, Norwood MA: Artech House, 2004. [7] Ehrfeld, W., Hessel, V., and Lowe, H., Microreactors: New Technology for Modern Chemistry, Weinheim, Germany: Wiley-VCH, 2000. [8] Quake, S., and Scherer, A., “From Micro to Nanofabrication with Soft Materials,” Science, Vol. 290, 2000, p. 1536. [9] Uriarte, L., Herrero, A., Ivanov, A., Oosterling, H., Staemmler, L., Tang, P., and Allen, D., “Comparison between Microfabrication Technologies for Metal Tooling,” Proceedings of the Institution of Mechanical Engineers, Part C: Journal of Mechanical Engineering Science, Vol. 220, 2006, p. 1665. [10] Lim, H., Wong, Y., Rahman, M., and Edwin Lee, M., “A Study on the Machining of HighAspect Ratio Micro-Structures Using Micro-EDM,” Journal of Materials Processing Technology, Vol. 140, 2003, p. 318. [11] Ehrfeld, W., and Lehr, H., “Deep X-Ray Lithography for the Production of ThreeDimensional Microstructures from Metals, Polymers and Ceramics,” Radiation Physics and Chemistry, Vol. 45, 1995, p. 349. [12] Chichkov, B., Momma, C., Nolte, S., Von Alvensleben, F., and T "unnermann, A., “Femtosecond, Picosecond and Nanosecond Laser Ablation of Solids,” Applied Physics A: Materials Science & Processing, Vol. 63, 1996, p. 109. [13] Gomez, D., Goenaga, I., Lizuain, I., and Ozaita, M., “Femtosecond Laser Ablation for Microfluidics,” Optical Engineering, Vol. 44, 2005, p. 051105. [14] Ross, R., Handbook of Metal Treatments and Testing, London, UK: Chapman and Hall, 1988. [15] Schubert, K., Brandner, J., Fichtner, M., Linder, G., Schygulla, U., and Wenka, A., “Microstructure Devices for Applications in Thermal and Chemical Process Engineering,” Nanoscale and Microscale Thermophysical Engineering, Vol. 5, 2001, p. 17. [16] Becker, H., and Gärtner, C., “Polymer Microfabrication Methods for Microfluidic Analytical Applications,” Electrophoresis, Vol. 21, 2000, p. 12.


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[17] Muhlberger, H., Hwang, W., Guber, A., Saile, V., and Hoffmann, W., “Polymer Labon-a-Chip System with Electrical Detection,” IEEE Sensors Journal, Vol. 8, 2008, p. 572. [18] Martinez, A., Phillips, S., and Whitesides, G., “Three-Dimensional Microfluidic Devices Fabricated in Layered Paper and Tape,” Proceedings of the National Academy of Sciences, Vol. 105, 2008, p. 19606. [19] Sure, A., Dillon, T., Murakowski, J., Lin, C., Pustai, D., and Prather, D., “Fabrication and Characterization of Three-Dimensional Silicon Tapers,” Optics Express, Vol. 11, 2003, p. 3555. [20] Daly, D., Stevens, R., Hutley, M., and Davies, N., “The Manufacture of Microlenses by Melting Photoresist,” Measurement Science and Technology, Vol. 1, 1990, p. 759. [21] Xia, Y., and Whitesides, G., “Soft Lithography,” Annual Review of Materials Science, Vol. 28, 1998, p. 153. [22] Chou, S., Krauss, P., and Renstrom, P., “Nanoimprint Lithography,” Journal of Vacuum Science and Technology B, Vol. 14, 1996, p. 4129. [23] Bourlon, B., Wong, J., Mikó, C., Forró, L., and Bockrath, M., “A Nanoscale Probe for Fluidic and Ionic Transport,” Nature Nanotechnology, Vol. 2, 2007, p. 104. [24] Terekhov, D., and O’Meara, M., “Recycling Metals Using the MOCVD Process,” Fourth International Symposium on Recycling of Metals and Engineered Materials, 2000, p. 487. [25] Leskela, M., and Ritala, M., “Atomic Layer Deposition (ALD): From Precursors to Thin Film Structures,” Thin Solid Films, Vol. 409, 2002, p. 138. [26] Williams, K., Gupta, K., and Wasilik, M., “Etch Rates for Micromachining Processing— Part II,” Journal of Microelectromechanical Systems, Vol. 12, 2003, p. 761. [27] Bu, M., Melvin, T., Ensell, G., Wilkinson, J., and Evans, A., “A New Masking Technology for Deep Glass Etching and Its Microfluidic Application,” Sensors and Actuators A: Physical, Vol. 115, 2004, p. 476. [28] Li, X., Abe, T., and Esashi, M., “Deep Reactive Ion Etching of Pyrex Glass Using SF6 Plasma,” Sensors and Actuators A: Physical, Vol. 87, 2001, p. 139. [29] Bourouina, T., Masuzawa, T., and Fujita, H., “The MEMSnas Process: Microloading Effect for Micromachining 3-D Structures of Nearly All Shapes,” Journal of Microelectromechanical Systems, Vol. 13, 2004, p. 190. [30] Chou, T., and Najafi, K., “Fabrication of out-of-Plane Curved Surfaces in Si by Utilizing RIE Lag,” The Fifteenth IEEE International Conference on Micro Electro Mechanical Systems, 2002, p. 145. [31] Jia, C., Wiemer, M., and Gessner, T., “Direct Bonding with on-Wafer Metal Interconnections,” Microsystem Technologies, Vol. 12, 2006, p. 391. [32] Burns, M., Johnson, B., Brahmasandra, S., Handique, K., Webster, J., Krishnan, M., Sammarco, T., Man, P., Jones, D., Heldsinger, D., et al., “An Integrated Nanoliter Dna Analysis Device,” Science, Vol. 282, 1998, p. 484.

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[33] Laser, D., and Santiago, J., “A Review of Micropumps,” Journal of Micromechanics and Microengineering, Vol. 14, 2004, p. R35. [34] Unger, M., Chou, H., Thorsen, T., Scherer, A., and Quake, S., “Monolithic Microfabricated Valves and Pumps by Multilayer Soft Lithography,” Science, Vol. 288, 2000, p. 113. [35] Meng, E., and Tai, Y., “A Parylene MEMS Flow Sensing Array,” Proceedings of the Transducers 2003 Conference, 2003, p. 686. [36] Berthet, H., Jundt, J., Durivault, J., Mercier, B., and Angelescu, D., “Time-of-Flight Thermal Flowrate Sensor for Lab-on-Chip Applications,” Lab on a Chip, 2010, p. DOI: 10.1039/c0lc00229a. [37] Catanzaro, B., Gillett, D., Simmons, M., Fennelly, J., and Sage Jr, B., “High Accuracy Non-Contact Optical Flow Sensor for Monitoring Drug Delivery,” SPIE Conference Series, Vol. 5691, 2005, p. 85. [38] Venkatasubramanian, R., Siivola, E., Colpitts, T., and O’quinn, B., “Thin-Film Thermoelectric Devices with High Room-Temperature Figures of Merit,” Nature, Vol. 413, 2001, p. 597. [39] Snyder, G., Lim, J., Huang, C., and Fleurial, J., “Thermoelectric Microdevice Fabricated by a MEMS-like Electrochemical Process,” Nature Materials, Vol. 2, 2003, p. 528. [40] Tang, W., Nguyen, T., and Howe, R., “Laterally Driven Polysilicon Resonant Microstructures,” Sensors and Actuators, Vol. 20, 1989, p. 25. [41] Johnson, W., and Warne, L., “Electrophysics of Micromechanical Comb Actuators,” Journal of Microelectromechanical Systems, Vol. 4, 1995, p. 49. [42] Kondo, R., Takimoto, S., Suzuki, K., and Sugiyama, S., “High Aspect Ratio Electrostatic Micro Actuators Using LIGA Process,” Microsystem Technologies, Vol. 6, 2000, p. 218. [43] Eaton, W., and Smith, J., “Micromachined Pressure Sensors: Review and Recent Developments,” Smart Materials and Structures, Vol. 6, 1997, p. 530. [44] Chen, P., Rodger, D., Agrawal, R., Saati, S., Meng, E., Varma, R., Humayun, M., and Tai, Y., “Implantable Micromechanical Parylene-Based Pressure Sensors for Unpowered Intraocular Pressure Sensing,” Journal of Micromechanics and Microengineering, Vol. 17, 2007, p. 1931. [45] MicroBUILDER Brochure, http://www.thinxxs.com/fileadmin/ website/pdf/MicroBUILDER_brochure.pdf. [46] Murphy, E., Inoue, T., Sahoo, H., Zaborenko, N., and Jensen, K., “Solder-Based Chipto-Tube and Chip-to-Chip Packaging for Microfluidic Devices,” Lab on a Chip, Vol. 7, 2007, p. 1309. [47] Peles, Y., Srikar, V., Harrison, T., Protz, C., Mracek, A., and Spearing, S., “Fluidic Packaging of Microengine and Microrocket Devices for High-Pressure and HighTemperature Operation,” Journal of Microelectromechanical Systems, Vol. 13, 2004, p. 31. [48] Link, D., Grasland-Mongrain, E., Duri, A., Sarrazin, F., Cheng, Z., Cristobal, G.,


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2 Microfluidic Building Blocks

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2.1 Introduction Microfluidics has become an enormous field. Multiple fabrication technologies, hundreds of flow control, separation and sensing techniques, and emerging applications in fields ranging from biotechnology to the chemical and oil industries all combine together to confuse and intimidate the scientist, engineer, or student approaching this area. To further complicate things, microfluidics is a field that is rapidly evolving: many publications dating from only a few years ago describe measurement techniques or manufacturing technologies that have since been superseded by newer, or more efficient, technologies. This chapter intends to be a selective guide to the different functional blocks that are available to the engineer for designing single phase and multiphase microfluidic devices, and for incorporating sensing elements in their designs to create a fully contained system. Since this book’s main focus is microfluidic integration, attention will be given to the fabrication and actuation technologies used in manufacturing different microfluidic subsystems, and their relative advantages and disadvantages will be outlined. The review presented here is very selective, and by all means is far from complete; examples and figures were chosen solely for their ability to reflect certain points made in the text, and these choices do not reflect any preference, endorsement, or guarantee. Section (2.5) is short compared to the others. This is a consequence of the fact that often detection technologies are very application-specific, whereas fluid manipulation technologies (either single-phase or multiphase) are required in all microfluidic applications. References will be given to review articles where possible for broader overviews of the subject matter. In addition to the publications and review articles cited in the text, a number of reference books [1–5] treat different aspects related to this chapter’s focus. The reader is directed to them for alternative (and, often, more detailed) approaches to the subject matter. References will also be made to some notions developed in Chapter 3, which covers the main physical phenomena that are dominant at the microscale. The reader should consult Chapter 3 for explanations regarding the physical principles governing the following devices.

2.2 Functional Approach Like many engineering problems, the development of a new microfluidic sensor can be performed using a functional block approach: the design is divided into several individual bricks that are each sized and optimized according to their function and specifications, and then manufactured using specific technologies and connected together to form a complete


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system. A good example is the development of a new car engine: the engine block, pistons, valves, spark plugs, fuel and oil pumps, sensors, and control systems are engineered and manufactured separately (often by factories on different continents) and they all work together very well as a system, although they rely on very different fabrication technologies (casting, machining, injection molding, semiconductor microfabrication). Unlike many engineering problems, however, microfluidics deals with flow at the microscale, in systems that involve channels of only a few micrometers in diameter and that contain and perform complex analyses on tiny fractions of a fluid drop. At these scales, issues related, for example, to dead volumes at fluidic connections, to dispersion of reactants in channels under laminar flow, and to capillarity, become extremely important, to the point where they can degrade the performance of the system to such an extent that it becomes useless. Often, therefore, microfluidic systems cannot be designed by combining individually manufactured functional blocks into a complete design. The constraints in choosing the individual blocks become stricter for fluidic systems: not only does each individual block need to perform its task well, but they also need to be capable of integration within a specific technology paradigm. Therefore, function and technology become intimately connected in microfluidics, and one needs to clearly define which tasks can be externalized and which need to be grouped together in a monolithic system. In the following we will pay particular attention to these aspects when reviewing systems capable to perform one or several fluidic control functions. We will include in this chapter different microfluidic blocks that perform individual flow control or sensing and measurement operations. The remainder of this book will then be dedicated to learning how to create a microfluidic design and how to analyze and optimize different microfluidic components. Several examples of commercial highly integrated systems will be given in Chapter 5.

2.3 Single-Phase Fluid Manipulation Some of the most important functions that need to be accomplished in a microfluidic system relate to fluid manipulation. The fluid, initially in a recipient external to the system (which might be a reservoir, a capsule, an industrial equipment, or a human body), needs to be guided to the right place within the microchip. Other times, fluids initially located within the microchip need to be pumped out, such as to deliver very small amounts of liquid (which might be a medication or a chemical reagent) in a controllable way. To accomplish these functions, one may need to actuate valves and pumps. The fluid may need to be mixed with another fluid, to dilute a sample or to perform a controlled chemical reaction. Since mixing in microfluidic

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Figure 2.1 Schematic view of a normally open microvalve using an external solenoid actuating an external plunger (left) and, respectively, of a normally closed microvalve with an integrated, magnetically actuated, metallic ball valve (right).

systems is not always a trivial issue, one may need to incorporate specific mixing mechanisms, which can be either active or passive. In this section we will review a number of the most commonly used techniques for flow manipulation. Excellent review articles exist on flow control technologies, and the reader is directed to those for more details [6–11]. 2.3.1 Active Valves Active valves rely on an actuation mechanism to allow, stop, or control flow within a microfluidic system. Very diverse systems have been imagined and realized, utilizing a wide array of actuation mechanisms. Below we will outline the most important mechanisms used, the devices realized, and the technologies used for fabrication. 2.3.1.1

Electromagnetic Actuation

Historically, magnetic actuation using external solenoids has been the first technology used for microvalve actuation [12]. It relies on the interaction of a magnetic material or coil with an external field produced by a separate solenoid. The magnetic material may be shaped as an external plunger or armature that applies local force to, for example, a microfabricated silicon [12] (Figure 2.1, left panel) or elastomeric membrane [13] (Figure 2.2), or it may be integrated within the microchip in the form of a magnetic layer (made of materials such as Permalloy, a widely used magnetic Ni-Fe alloy that can easily be fabricated by electroplating at or near room temperature) deposited


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on a mobile structure [6, 11, 14], or in the form of a metallic ball [15]. Figure 2.1 (right panel) shows schematically a possible implementation of a normally closed ball valve using an external electromagnet for actuation. The relative advantages and disadvantages of using external solenoids are discussed below and summarized in Table 2.1. The main advantage is that relatively large magnetic fields can be switched on and off at will, which can actuate several magnetic valves simultaneously, and that only relatively simple manufacturing processes are required on the microfabrication side. There are a number of drawbacks, however. Due to the significant size of the required external components (in particular, the electromagnets’ solenoid coils and cores), the lateral dimensions of the final device are necessarily large (several millimeters on each side being required per actuator), preventing dense packing of individually addressable components. The electromagnetic actuation requires significant power consumption. The fabrication process is likely not compatible with batch fabrication (e.g., whenever manual placement of magnetic balls within valve cavities is required), hence increasing cost. While the magnetic fields generated using external solenoids can be significant, the resulting force (and corresponding maximum valve differential pressure) depends on the amount of magnetic material present in the microsystem: metallic bead microvalves, for example, can operate against much higher pressures (order of magnitude: 10 bar) than microvalves using deposited magnetic films (order of magnitude: 1 bar). The drawback in that case is that significant swept volumes are associated with ball microvalves: as an estimate, a ball of diameter 3 mm (typical size for this application) that is located in a cylindrical cavity of diameter 3.1 mm and height 3.5 mm, will result in a swept volume of approximately 12 µl, which is really high for a microfluidic application; in addition, a significant amount of this volume is dead volume, communicating only by diffusion with the rest of the flowing fluid (e.g., the fluid trapped in the corners of the cylindrical cavity). At an estimated flow rate of 1 µl/min, such a system will require approximately half an hour of cleanup time. As an alternative to ball valves, drops of hydrophobic ferrofluids (realized, for example, using a suspension of ferromagnetic particles in a hydrophobic fluorocarbon carrier) have been manipulated using external magnets, and used to block or allow flow through a channel [16]. Operation up to back pressures of ≈0.1 bar was recorded, with dead volumes that are much smaller than in the case of a solid ball valve. Commercial plunger-type solenoid valves exist, which are manufactured using nonlithographic technologies such as high-speed micromachining and precision injection molding (the LEE Company manufactures a large array of models for inkjet printing applications; see Figure 2.3). Such valves, which

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Figure 2.2 A PDMS system making use of the verstility of PDMS fabrication and incorporating external solenoid valves alongside screw and pneumatic valves. (Reproduced from [13] by permission of The Royal Society of Chemistry.)

Figure 2.3 Examples of commercial solenoid microvalves fabricated by the LEE Company. (Used with permission from the Lee Company, www.theleeco.com.)

can have internal volumes as low as a few microliters, can often be used instead of integrated microvalves in a variety of applications. Other types of magnetic actuation systems have been implemented using integrated microcoils [17]. In this case, a spatially varying external magnetic


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Table 2.1 External Solenoid Actuation with Internal Magnetic Material

Advantages Electrical on-off switching Simple microfabrication Strong magnetic fields

Disadvantages Significant power consumption Sparse integration of components Large total system volume Not a batch-process Important internal volume (ball valve) High cost

field is provided by using, for example, a permanent magnet placed outside or by soft magnetic layers integrated within the chip. Actuation is realized by energizing the coil, which results in the creation of a magnetic dipole. When placed in a spatially varying external magnetic field perpendicular to the plane of the coil, the coil will witness an electromagnetic force that is proportional to the product between the magnetic dipole moment (and hence the coil currents) and the value of the field gradient. Typical forces that can be achieved by this technique are on the order of 1 mN, and corresponding operating pressures (considering a seat of approximately 100 µm diameter) can be as high as 1 bar. The advantages of this technique consist of the ability to incorporate multiple independent on-off valves that are closely packed together, and with relatively low dead and swept volumes. Disadvantages stem from the relatively low operating pressures, complicated microfabrication with very specific, nonstandard processes (such as mold electroplating, and so on), power consumption, and side effects such as Joule heating of the microsystem. Cost may also be a concern, especially since for compact microfluidic applications very strong and compact permanent magnets are often required, such as rare-earth magnets (e.g., samarium cobalt and neodymium iron boron). These advantages and advantages are summarized in Table 2.2. 2.3.1.2 Piezoelectric

Piezoelectricity is the ability of certain crystalline materials to produce mechanical stress or to stretch as a consequence of being exposed to an electric field. Piezoelectric actuation is interesting in microfluidic systems as it provides a combination of two advantageous features: high force and low stroke, which can be combined together to enable control and efficient delivery of very small amounts of fluid. Consequently, many examples of systems

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Table 2.2 Internal Microcoil Actuation with External Magnet

Advantages Electrical on-off switching High component integration density Small internal volume Batch process

Disadvantages Significant power consumption Low operating pressures Specific microfabrication required Important Joule heating Cost

using piezoelectric actuators for pumping purposes exist commercially or have been published in the literature (see Section 2.3.3). Surprisingly, fewer examples of microvalves using piezoelectric actuation principles have been reported, one reason for this being precisely the low stroke that is achievable using piezoelectric structures: sufficient displacement of a valve seat to have a completely open valve can only be achieved by using stroke amplification mechanisms, and result in relatively complicated implementations. The most common stroke amplification scheme that has been employed is that of hydraulic amplification: a piezoelectric crystal actuates a piston of area S1 , which applies pressure to a high-modulus liquid trapped within a closed cavity. The pressure is transmitted to a second piston of lower surface area S2 , which is displaced by a distance d proportionally higher than the stroke of the piezoelectric crystal δ: d = δ SS12 . The realization of the sealed hydraulic cavity with pistons of different diameters is complicated, and the whole system requires up to nine different silicon, SOI, and glass layers bonded together and manually fitted with piezoelectric crystals and then filled with silicone oil [18]. The final system, fabricated at the Massachusetts Institute of Technology and representing a real tour de force in microfabrication and microassembly, is capable of providing a controlled displacement of the sealing membrane by up to 20 µm, resulting in a proportional valve that can operate against pressures of several bar and at frequencies up to several kilohertz. A schematic of the system is shown in Figure 2.4. Other types of microvalves incorporating piezoelectric actuators have been demonstrated at the Jet Propulsion Laboratory [19], utilizing much larger, external piezo stacks that achieve displacements of up to 5 µm. These normally closed valves have been used in micropropulsion applications, and have been able to operate against pressures of over 50 bar with very small leak rates in the closed position. In alternate implementations, piezoelectric bimorph actuators have been used to open and close microfabricated silicon [20] or plastic [21]


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Figure 2.4 An example of hydraulically amplified piezoelectric valve showing the schematic of the system (left) and a photograph of the realized microsystem (right). The main components of the valve are: 14— the piezoelectric crystal, 10—the large-diameter piston, 5—the small diameter piston and sealing element, 11—the oil-filled cavity, and 4—the valve orifice. The footprint is 2 cm on the side by 1 cm tall. (Reproduced from [18].)

Table 2.3 Piezoelectric Actuation of Valves

Advantages Electrical on-off switching High pressure operation Low power consumption

Disadvantages Complicated fabrication process Not a batch-process High voltage electronics Relatively large size High cost

valves to be used in drug delivery, and respectively, ink dosage systems such as fountain pens. Such actuators allow only on/off operation of the valve, proportional operation being achieved by pulse-width modulation techniques and requiring constant power consumption in this operation mode. Relative advantages and disadvantages of piezoelectric actuation are listed in Table 2.3. 2.3.1.3 Pneumatic / Thermopneumatic

Pneumatic actuation is widely used in microfluidics, allowing a high degree of on-chip integration. Pneumatic actuation involves the displacement of a flexible membrane under the application of external pressure from a

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Figure 2.5 Left: Schematic of pneumatic microvalve manufactured using multilayer soft lithography. The device consists of a fluidic channel and, perpendicular to it, a pneumatic or hydraulic channel, the two being separated by a thin flexible membrane (A). By pressurizing the pneumatic channel, the membrane is deflected, thus blocking the fluidic channel (B). Once pressure is removed, the elastic membrane regains its equilibrium shape (C). Right: Pneumatic microvalve network manufactured in PDMS, allowing complex manipulations to be performed on several fluid streams: 1—loading, 2—compartmentalization, 3—mixing, and 4—purging. The streams have been intentionally dyed to allow visualization of the different processes. (Reproduced with permission from [22].)

pressurized gas line. The membrane, placed between the gas line and a fluidic channel, bends under the pressure difference and partially or completely obstructs the flow through the channel. Thermopneumatic actuation, by contrast, operates under the same principle but uses pressure generated internally by, for example, the thermal expansion of a heated fluid, by changing its vapor pressure or by thermally nucleating a gas bubble within a closed cavity. The most impressive results using pneumatic actuation have been achieved by Quake et al. [22, 23] using a silicone elastomer (PDMS) as the manufacturing material and bonding multiple micromolded PDMS layers together to create a complex geometry (multilayer soft lithography). Crossed microfluidic and pneumatic channels were separated by a thin PDMS membrane, which deflected and pinched off the fluidic channel when gas pressure was applied to the pneumatic channel. The pneumatic lines were controlled using external solenoid valves. This technique proved to be extremely versatile, allowing the simple fabrication of extremely small dead volume valves with different actuation pressures (by changing the lateral channel dimensions), and enabling complex fluidic manipulations (Figure


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Table 2.4 Pneumatic Actuation of Valves

Advantages Rapid on-off switching Dense integration High versatility Simple fabrication

Disadvantages High power consumption Large external equipment High cost

2.5). In addition to their use as valves, several such structures can be actuated sequentially in a peristaltic regime to generate flow; highly integrated systems using thousands of microvalves and micropumps operated in parallel have been realized with this technology. A simple model of this type of valve is analyzed in Chapter 4, whereas commercial applications of this technology are further described in Chapter 5. Several variations of this technique have been developed, using, for example, external pneumatic tubes instead of integrated microchannels for the pneumatic lines [13] (Figure 2.2). While cross-channel configurations work well with highly flexible, low Young modulus materials such as PDMS, when using more rigid materials, one has to revert to a more traditional membrane microvalve geometry [24, 25]. An interesting twist is provided by Luque et al. [26], who used a combination of surface and bulk micromachining to create a monolythical microvalve. Two such microvalves, when placed back to back, allow very convenient operation as a pneumatic valve, where the membrane displacement is automatically limited by the device. This configuration, shown schematically in Figure 2.6, allows high-pressure operation in addition to relatively simple and economical fabrication. Recently, the use of integrated inert plastic materials (PEEK) as microvalve membrane materials within natural gas analyzers has been demonstrated [27]. Table 2.4 lists the relative advantages and disadvantages of pneumatic microvalves. Thermopneumatic actuation uses the variation of a solvent’s vapor pressure with temperature or the thermal dilation of a liquid or gas to apply pressure to the valve diaphragm. Using a highly volatile solvent like pentane and built-in heaters to rapidly change its temperature allows relatively high actuation pressures to be generated, assuring the closing of the valve against pressures as high as 1.3 bar [28]. Such a thermopneumatic system, shown in Figure 2.7, is naturally sensitive to ambient temperature variations, and requires careful thermal analysis: unless capable of efficient cooling, such a system risks to have slow on-off cycles on the order of several seconds, due

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Figure 2.6 Schematic showing two thin-flim membrane valves placed back to back, in a configuration allowing convenient high-pressure pneumatic operation and simple fabrication. (Adapted from [26], with permission from Elsevier.)

Figure 2.7 Thermopneumatic microvalve using the change in vapor pressure of heated pentane as an actuation mechanism. System realized using silicon/glass micromachining and a corrugated silicon membrane as the valve’s moving structure. (Reproduced from [28].)

to the slow heating and cooling of the entire sensor structure. This issue has been recognized by other research groups, who managed to partially avoid


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Table 2.5 Thermopneumatic Actuation of Valves

Advantages Complete on-chip integration Low cost Large stroke

Disadvantages High power consumption Very slow operation Complicated fabrication / filling Thermal cross-talk Low pressure

Figure 2.8 Electrostatic valve design that makes use of pressure balancing and of a flexible electrode to improve both the operating pressure and the openposition flow gap of an electrostatic microvalve. (Reproduced from [30], with permission from Elsevier.)

it by integrating water cooling channels within the chip [29] (in that study, heated air was used as the actuation medium and a silicone elastomer was used to define the valve membrane). Valve densities of the order 300 cm−2 could be achieved this way, with very little thermal cross-talk. 2.3.1.4 Electrostatic

Electrostatic actuation relies on the electrostatic attractive force that develops between the valve membrane and the valve seat when a voltage difference is applied. Due to electrolysis issues in conductive fluids, electrostatic actuation has been primarily used with gas valves—while their direct applicability to microfluidics is currently limited, future designs that eliminate electrolysis issues may be possible. In addition, electrostatic gas microvalves may be used as pilot valves for fluidic valves operating against higher pressures. Therefore we include a brief review of some of the results in this area.

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Table 2.6 Electrostatic Actuation of Valves

Advantages Complete on-chip integration Low power consumption Fast operation

Disadvantages Limited to gas applications Small gap when open Complicated multi-layer fabrication Low pressure

Electrostatic actuation is very effective for small gaps: the force between the two structures is inversely proportional to the square of the gap between them (considering constant voltage operation), and thereby rapidly decreases with distance. Because of this issue, operating pressures of electrostatic valves tend to be relatively small (a fraction of a bar), and the flow gaps when the valve is in the open position are limited to a couple of micrometers. To overcome the limited range of electrostatic actuation mechanisms, valve strokes have been increased by combining electrostatic actuation with other types of opening and closing mechanisms (e.g., using the Lorentz force between permanent magnets and a conductor placed on the valve membrane), and the electrostatic force was only used to keep the valve closed with minimum power consumption [31]. Alternatively, compliant diaphragms have been used in conjunction with pressure balancing mechanisms to increase both the operating pressure and the gap in the open position [30]. The resulting system, shown in Figure 2.8, is in principle capable of opening at a voltage of 10V against 1 bar of inlet pressure, while maintaining a 5 µm gap in the open position. In another study, electrostatic microvalves have been used to control the air flow within pneumatic tactile Braille reading systems. Prototype valves using a polysilicon moving element have been demonstrated to open against 1 bar of differential pressure at an actuation voltage of 92V RMS [32]. An elegant example of electrostatic microvalve is represented by the fuel injection valve of the MIT microengine project [18, 33]. The very ambitious design utilizes a stack of three fusion-bonded Si and SOI doped wafers and involves a planar moving electrode as a sealing structure. The device was pressure balanced and was capable of opening against 9 bar differential pressure with an actuation voltage of 136V. It is represented in Figure 2.9. 2.3.1.5

Latching Valves

Active valves generally require continuous actuation during the open state (for normally closed valves) or the closed state (for normally open valves).


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Figure 2.9 The electrostatic valve designed for the MIT microengine project involves a complicated structure of SOI wafers machined using DRIE techniques and fusion bonded. (Reproduced from [18].)

This can lead to large power consumption and to difficulties in simultaneous multiple valve addressing. Latching valves eliminate these inconveniences by requiring actuation power only to switch from one state to the other. After having switched, the valve stays in that state until the next time it is actuated. This type of bistable operation, which can be considered a form of valve memory, can be extremely beneficial in many applications, particularly where multiple valves need to be controlled independently. Several types of latching valves have been demonstrated in microfluidic systems. In the simplest configuration, an external bistable electromagnetic actuator was coupled to a microfluidic valve consisting of a thin elastomeric membrane bonded between two micromachined silicon wafers [34] (Figure 2.10). The valve is expected to be particle tolerant due to its soft elastomeric membrane. The bistable action was not intrinsic to the microfluidic system, but was rather provided by the actuator itself. Due to the external actuators, this latching valve configuration is not well suited for integration—the density of valves that can be included in a design is severely limited. A significantly more complex design, but this time integrating all the required components within the microfluidic system, was manufactured using silicon microfabrication and manual assembly and bonding between the parts [35]. The valve consisted of a fixed microfabricated gold coil and a moving cantilever beam covered with a magnetized foil. The coil was used to actuate the beam between its two stable positions, where it was being held by a combination of elastic and magnetic forces. The valve was designed to open against 2,000 Pa and was tested for opening and closing in both air and water;

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Figure 2.10 Example of an electromagnetic latching valve using an external bistable actuator. Left: Schematic of the valve in the closed (top), and open (bottom) position. Right: Three-dimensional representation of the valve. The external actuator moves the closing boss by approximately 200 µm. (Reproduced from [34], with permission from Elsevier.)

however, no other valve operating parameters were published. Due to its hard seat, this valve is not expected to have any particle tolerance. The manual assembly steps using adhesive bonding make this technology unsuitable for batch fabrication in its current form. An original approach to creating a latching valve involves the phase change of a wax material from the solid to the liquid phase upon heating, and back to the solid phase upon cooling [36]. When the heaters were actuated, the wax droplet melted and could be pushed into and out of a fluid channel by using an external pressure or vacuum source. Once the power to the heaters was stopped, the wax solidified in its current position, either blocking the fluid channel or leaving it open. This valve design is particle tolerant (any particles being incorporated in the wax when melted). A fully integrated microfluidic design with pneumatic valve actuation was manufactured from two machined glass wafers bonded using a thin elastomeric membrane [37] (Figure 2.11). The operation depended upon the normally closed nature of the valves, latching valve configurations consisting of both three- and four-valve circuits having been demonstrated. Vacuum or pressure pulses as short as 120 ms were sufficient to switch the latching valves between the open and closed positions and to hold them in that state for tens of seconds. Even more impressively, such latching valves can be used to create a multiplexor, thus allowing up to 2n latching valves to be actuated independently using only n pressure/vacuum lines.


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Figure 2.11 A latching valve design involving pneumatically actuated valves. The input channel is used to send either a vacuum pulse, or a pressure pulse, which switches the states of the vacuum and pressure valve temporarily, causing a permanent change of the latching valve state. (Reproduced from [37] by permission of The Royal Society of Chemistry.)

2.3.2 Passive Valves Contrary to active valves, which involve an actuator and can be operated on demand, passive valves represent mechanical elements within a microfluidic system that open in one direction under the action of pressure to allow the flow of fluid and stay closed under reverse pressure conditions, but cannot be controlled externally. Such valves can be used in microsystems as isolation elements, separating the internal fluid from the external one and thus preventing sample contamination or evaporation over long periods of time. Another common application of microfluidic passive valves is in the realization of reciprocating and peristaltic pumps, where they are used at the inlet and outlet to ensure that fluid always flows in the right direction. In both applications, leak rates are important and need to be minimized, since a leaky check valve means a less efficient pump. Another parameter that is important is the opening pressure of the valve, which needs to be minimized as well (unless the valve needs to be used to generate back pressure), as well as the maximum opening gap—normally the valve includes a stopper that limits the travel, thus preventing failure for high forward pressure. The maximum reverse pressure that a valve can withstand is also a crucial parameter, since it determines an important failure mode and limits the applications in which the valve can be used. Finally, for applications related to high-frequency pumping, the resonant frequency of the valve is important and ideally should be high to allow rapid operation. This implies stiffer valves involving less mobile mass; however, there are designs where the actuation of the pump can be performed at frequencies above the natural resonant frequency of its valves. The choice of

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Figure 2.12 Passive valve design schematics of a membrane-type (left) and a flaptype check valve device (right)

.

a specific design depends on the application at hand, which dictates the ranges of parameters within which the valve needs to operate. As an example, in space applications such as micropropulsion, the requirements for microvalves are: (1) the ability to support pressures of 10 MPa or larger, (2) the flow rate of 10 cm3 /s or larger, (3) the operating frequency of 10 kHz or larger, and (4) low or no power consumption. 2.3.2.1

Check Valves

Many examples of microfluidic check valves exist in the literature. Typically, these involve some sort of elastic element such as a membrane or a flap (Figure 2.12), which allows the passage of the fluid in one direction but seals against the orifice when pressure is applied in reverse. Perhaps one of the oldest designs of a microfluidic passive valve manufactured using MEMS techniques and having an elastic membrane as the mobile element was incorporated in micropumps developed at Twente University in the Netherlands in the late 1980s [38, 39]; similar designs have been used numerous times since then. The most common design involves a membrane connected by a number of radial bridges or an elastic flap, which act, on one hand, as elastic elements and, on the other hand, allow the fluid to pass when the valve opens [6]. The development of passive valves has been mostly governed by applications to pumping devices; they will therefore be discussed in more detail in Section 2.3.3. Fabrication of check valves is relatively simple and can be realized in only a few layers; they can be manufactured in a very wide range of materials. While the majority of published examples are manufactured in monocrystalline silicon [40–44] that has exceptional elastic properties, other materials that can be used are plastics or printed circuit board [45–48] (either machined, laser-cut, or precision injection-molded), photoresists [49] (such


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Figure 2.13 A thermoelectrically actuated microreservoir for drug delivery: when the reservoir needs to be opened, a voltage is applied across the metallic isolation membrane, which ruptures, thus allowing the drug to diffuse out of the reservoir.)

as SU8), elastomers [50] (such as PDMS), and thin metal films [51–54], electroplated or otherwise deposited. 2.3.2.2 Burst and Controlled-Release Valves

Another example of (typically) passive microfluidic valve is the burst or controlled release valve, which dissolves or ruptures and thus opens a certain portion of a microfluidic system when certain conditions are met—usually when the differential pressure reaches a threshold value, when the chemical environment is such that the isolation material of the burst valve dissolves, or when an electrical signal is sent that melts or otherwise destroys a separation membrane. Such valves are one-shot elements, and can be implemented using both passive and active technologies. Applications for this type of structure exist in, for example, drug delivery applications [55]. Systems using an electrochemical reaction that dissolves a metallic separation membrane have been developed [56] using a gold layer as the separation element and a saline solution as the medium. Other examples involve reservoirs capped with resorbable polymeric caps that may have different thicknesses and chemical compositions, and thus allow release at different time intervals or when the medium reaches certain critical chemical compositions [57]. An example utilizing active thermoelectric action has been developed by MicroCHIPS [58] for timed drug delivery. Using a metallic membrane, this valve allows simple electrical actuation—an electric pulse is sent across the membrane that heats by the Joule effect and ruptures (Figure 2.13). This technology has the advantage of being medium-independent;

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a

b Figure 2.14 Panel a: The Tesla valve, as described in the original 1916 patent, showing the casing (1), the tubing connection (4), and Tesla valve conduits meant to create asymmetry between forward and reverse flow (3). (Reproduced from [59].) Panel b: Computational fluid dynamics simulation of the flow velocity field within a microfluidic Tesla valve with channel width of 114 µm under reverse flow conditions, at a Reynolds number equal to 528. (Reproduced from [60], with permission from University of Washington.)

such membrane burst schemes could be applied in any application and do not require, as some of the passive examples described above, a conductive medium. 2.3.2.3

No-Moving-Parts Hydrodynamic Valves

One type of passive valve that is particularly easy to manufacture using a single etch procedure is the hydrodynamic valve, which involves no moving parts. This type of valve is always open (cannot sustain any amount of back pressure), and it is based on the asymmetry between forward and reverse flow in certain geometries (see Figure 2.14) under specific flow


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conditions [61]. In general, such asymmetry does not appear at the very low Reynolds numbers usually encountered in microfluidics, but starts to develop as the Reynolds number is increased (e.g., by increasing the flow rate). The parameter that quantifies the valve asymmetry is called diodicity (alluding to the similarity with an electrical diode) and is calculated as the ratio of pressure drops for equal magnitude in forward and reverse flow rate conditions. The hydrodynamic valve concept has been used in certain pumping applications [62, 63] (typical diodicity: 1.2), where the main advantages brought by this type of valve are the particle tolerance and the ease of manufacturing. Numerical analysis of such valve concepts suggests that in certain flow conditions diodicities of up to 1.7 are possible by valve design optimization [60]; however, it was shown that it is not possible to design a hydrodynamic valve capable to operate for a wide range of flow rate conditions [61]. The applications of this type of valve remain therefore quite limited in microfluidics. 2.3.3 Pumping Techniques One of the most important components of a microfluidic system is its pumping mechanism—fluids require a pressure difference to generate flow through the microfluidic channels, and the source of that pressure difference can be of diverse origins. Microsystems may generate that pressure difference internally, or they may make use of existing external pressure sources. This section will try to review a few of the most commonly used flow-generation techniques in microfluidics, and outline specific particularities, advantages, and trade-offs of each. 2.3.3.1 Externally Imposed Flow Rate

By far, the simplest situation (and also the most common, at least in the laboratory setting) is to use an external volumetric pump, typically a microprocessor-controlled syringe pump. There are many types of pumps and syringes available, which can be combined to reach almost any flow rate regime. Large syringe pump manufacturers include KD Scientific and Harvard Apparatus (the most common models are shown in Figure 2.15), which are now part of the same company but maintain separate product lines. The most usual design for syringe pumps is to use a screw actuated by a stepper motor; this provides very precise lateral displacement of the pump piston. However, due to the discrete motion of the stepper motor, the flow at the output of the syringe pump may be pulsed, which can be a disadvantage in certain applications. The usual way to avoid this is the use of compliant tubes and syringes, which act as a lowpass filter and dampen the fluctuations

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Figure 2.15 Examples of syringe pumps that are common in laboratory settings. Left: A model manufactured by Harvard Apparatus. Right: The equivalent model from KD Scientific. (Reproduced with permission from Harvard Apparatus.)

created by the pump, at the cost of increasing the time response of the system. Unfortunately, despite their universal use in the testing and calibration phase for microfluidic systems, syringe pumps tend to be large, heavy, expensive, and power-hungry, which limits their area of application to laboratory settings. 2.3.3.2

Pressure-Driven Flow

One way to completely remove the fluctuations in flow rate is to use pressuredriven flow instead of a syringe pump. In a laboratory setting, one can either use an external pressure source such as a pressurized gas cylinder or simply use the hydrostatic pressure of a column of water. In an industrial setting, most fluid conduits that may require monitoring using microfluidic sensors are already pressurized or develop important pressure drops along the piping. Such inherent pressure differences can be used to generate flow through the microfluidic portion of the system without the use of a dedicated pump (in this case the microfluidic system is used as a bypass). The same may be true of implantable systems used for measuring certain metabolic parameters—the pressure difference that develops between various points in the body (such as between an artery and the return vein, or between an artery and the external environment) can be used to passively drive blood through a microsystem. In order to control the flow in the case of pressure-driven systems, one requires some sort of active valve. Several technologies were discussed in Section 2.3.1 and can be used to this end. In the case of fluids containing particles or biological cells (blood being an example), a valve design that is tolerant


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to particles needs to be used. Typically, this involves manufacturing in a soft material such as an elastomer. 2.3.3.3 Centrifugally-Driven Flow

An original technique for driving fluids within microfluidic channel networks consists of using a rotating substrate, similar to a compact disc incorporating microfluidic channels. The centrifugal force generates a pressure gradient within the microfluidic channels, which moves fluid outwards. Since the rotational speed can be changed by several orders of magnitude, the driving force can be varied at will. The centrifugal platform typically implements only passive fluidic functions due to the difficulty of accessing different components while the disk is spinning. In this sense, it looks like a less versatile implementation of pressure-driven pumping. One feature, however, sets it apart and allows it to implement certain unique functions: being a rotating system, it can make use of the Coriolis force to implement rapid mixing [64] and fluid switching [65] techniques. 2.3.3.4 Mechanical Pumping / Reciprocating Motion

As mentioned already, research in the area of micropumps and microvalves was initiated in the late 1980s at Twente University by the works of van Lintel and Smits [38, 39, 66], who developed the first modern microfabricated reciprocating pumps (Figure 2.16). In the more than two decades that followed that development, hundreds of other designs and pump geometries have been published, of which several have evolved to the commercial stage. The next few sections will outline just a few examples to give the reader an idea of the diversity of techniques employed. For a more complete view, the reader is directed to some of the excellent review articles mentioned in the beginning of the chapter [7–11, 67]. One of the most common pumping technologies that has been applied to microfluidic systems is that of the reciprocating pump—the vast majority of published pump designs fit within this category. The operating principle relies on a combination of a variable volume chamber and of two passive check valves connected to the chamber in opposite fashion. When the volume of the chamber is increased (depression cycle), the input check valve opens and the output valve closes. In this configuration the valve fills with fluid from the inlet channel. When the volume of the chamber is then decreased (compression cycle), the input check valve closes and the output valve opens, pushing the fluid out of the valve body and into the outlet channel. The pump may use active elements for both depression and compression actions, or just for one of them. In that case, the opposite action is normally achieved by the

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Figure 2.16 The early piezoelectrically actuated micropump developed by van Lintel at the University of Twente. (Reproduced from [38], with permission from Elsevier.)

elastic relaxation of the pump cavity. Pumps have been designed that have a single pumping cavity or multiple ones disposed in series or parallel. The design and fabrication of passive check valves suitable for use in pumping applications are a difficult undertaking. Several important criteria such as presence and tolerated amount of backward flow, pressure drop under operation, switching speed, bubble and particle tolerance, and maximum reverse pressure capability need to be properly balanced to achieve a working micropump. The pump and valve materials need to be selected carefully— interactions and chemical compatibility with the pumped fluids, or simply materials’ fatigue due to millions of on-off cycles, can drastically change the properties of the materials and thus the pump operation. This is particularly true for polymeric pumps, and is less of an issue for monocrystalline silicon devices. The valves and other passages in a micropump can become clogged or obstructed by particles, which can completely compromise the pump operation—particle filters at the inlet or soft seals may need to be integrated to minimize the chance of failure (particle tolerance of more compliant polymeric pumps is normally better than that of pumps made of harder materials such as silicon). The bubble tolerance and the self-priming capability of the pump are important criteria as well. Since the presence of a bubble can drastically increase the effective compressibility of the pumped medium, it will correspondingly alter the pump efficiency; high compression ratio per stroke and airtight check valves are required for operation with gases and for improved bubble tolerance. These numerous criteria are just a few of the issues that need to be weighed in the process of designing and manufacturing a reliable micropump for a specific application. The large majority of published reciprocating micropump designs have been manufactured in hard materials such as silicon and glass—29


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Figure 2.17 The typical operational curve for a reciprocating pump (solid line). The intercepts are Φmax and δpmax , which are two of the important merit criteria used in pump evaluation. The dashed curve shows the work performed by the pump on the fluid per unit time (or pumping power). The operating point of a micropump to be used in a battery operated system is typically chosen to maximize the pumping power while minimizing electrical power consumption.

of the 46 micropump systems covered by an extensive review [7] were manufactured using variants of this technology. Other materials that have been used in micropump design include plastics (injection molded and/or laser-cut polycarbonate or cyclic olefin copolymer (COC)), elastomers, and metals (usually brass). Planar technologies are almost ubiquitously used in the fabrication of micropumps. The variable volume chambers are typically cavities covered with a flexible membrane. Actuation of the membrane can be achieved in many different ways, most of them similar to the active valve actuation techniques described in Section 2.3.1—external electromagnetic actuators (e.g., solenoid with plunger), disk or bimorph cantilever-type piezoelectric crystals, piezoelectric stacks, pneumatic and thermopneumatic actuators, and integrated electrostatic and electromagnetic actuators. Among these, piezoelectric actuators are by far the most common, thermopneumatic and pneumatic actuations being the second most commonly used technologies. Electrostatic actuation promises to become a viable actuation method. We invite the reader to review Section 2.3.1 for a detailed description of these actuation principles.

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There are several merit criteria that can be used for evaluating and comparing micropump performance. The maximum achievable flow rate under no differential pressure (Φmax ) is important in applications where the role of the micropump is to transport fluid as fast as possible through a system with relatively low hydrodynamic resistance. In general, Φmax is limited by the maximum frequency at which the pump elements can operate, and by the pump cavity volume. Another important parameter is the maximum achievable differential pressure under no flow conditions (δpmax ), which needs to be maximized in applications where the micropump is used to pressurize a system or to drive fluid through a high hydrodynamic resistance system or against an external pressure gradient (i.e., outlet at a higher hydrostatic pressure than the inlet). δpmax is typically limited by the maximum force that can be generated by the actuator, by valve sealing efficiency, and by the back-pressure tolerance of the inlet check valve. The energy efficiency of the pump η = Φδp/P , defined as the ratio between the work performed by the pump on the fluid per unit time (or pumping power, which can be quantified as the flow rate Φ times the pressure differential δp) and the electrical power consumed by the pump actuator P , is another parameter that is important for applications where high energy efficiency is required, such as in autonomous battery-operated systems. A typical reciprocating pump δp–Φ curve is shown in Figure 2.17; published energy efficiencies for reciprocating pumps are on the order of a fraction of a percent. The maximum pumping power Φδp, which can be approximated as 41 δpmax Φmax assuming Φ-δp linearity in Figure 2.17, can be used as a criterion for evaluating the overall pumping performace of the micropump. The size of the micropump can also be a deciding factor in the applicability to a specific use—generally, the trend in micropump design is towards smaller geometries, which are usually also more energy efficient. Size also affects cost (lowering or raising it depending on whether or not batch manufacturing is performed), which is important particularly for disposable pump applications. Finally, another important criterion for valve evaluation is the magnitude and reproducibility of the volume delivered per stroke, which is very important in microdosing and dispensing applications. It should be clear from the above that the diversity of available pump designs and actuation technologies, the wide choice of manufacturing materials, and the numerous distinct evaluation criteria can make the comparison of different micropumps very difficult if not impossible. Usually, a micropump needs to be designed and optimized for each specific class of applications, and then can only be compared against competing micropumps based on the criteria relevant to the application at hand. In the following, we will describe a number of reciprocating micropumps that show outstanding


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a

b Figure 2.18 (a) The piezoelectric microvalve developed at MIT uses a single pumping chamber and two passive check valves and (b) it involves a complex manufacturing process utilizing seven independent silicon, SOI, and glass layers bonded together, as well as a piezelectric cylinder as the actuation element. (Reproduced with permission from [68].)

performance as judged by some of the many evaluation criteria listed above. A selection of different manufacturing technologies was made to provide as wide an overview as possible. The high-pressure, high-flow rate, piezoelectrically driven micropump developed at MIT [68] has some of the best performance parameters among all published micropump designs. In the design of that micropump no shortcuts that could compromise performance were made: the best technology available at the time was used, with little regard for cost or manufacturing complexity. The microvalve, shown in Figure 2.18 both as a top-view schematic (Panel a) and in cross-section (Panel b), consists of seven independent layers of silicon, SOI, and glass wafers, bonded together using a combination of fusion, anodic,

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and eutectic bonding. The bonding steps involved a combination of waferlevel and package-level bonding. The actuation of the valve is performed using a piezoelectric cylinder 1 mm in diameter, which was eutectically bonded to the silicon and actively performed both the positive and the negative stroke actions; various layers in the structure were doped and acted as electrodes for actuation of the piezoelectric crystal. The device layers on the SOI wafers are used as the elastic pump and valve elements—the SOI technology allows very precise control over the thickness of the different planar components, since the buried oxide layer in the SOI can act as an etch stop for the deep reactive ion etching process. Overall, the manufacturing of the microvalve represented an impressive tour de force and the results were quite impressive: the valve achieved a maximum flow rate of 2 ml/min at a drive frequency of 4.3 kHz, in a 16×17×2 mm3 package (the active components of the device occupied only a quarter of the area of the device, the rest being devoted to electrical and fluidic connections). The operational curve for an actuation frequency of 3.5 kHz showed intercepts of δpmax = 300 kPa and Φmax = 1.3 ml/min, leading to a pumping power 14 δpmax Φmax ≈ 1.6 mW. The piezoelectric crystal required actuation with voltage pulses as high as 1,200V. Interestingly, the performance of the pump (as measured by the maximum achievable flow rate at zero differential pressure) as a function of frequency showed peculiar behavior: at certain frequencies the output of the pump dropped to zero, which is characteristic of resonances developing in the device. The observed resonances could not be recovered in finite element simulations so the authors could only speculate on their origin or on possible improvements or optimizations that might eliminate this peculiar behavior; this indicates again the complexity of designing a reliable, general-purpose micropump. The manufacturing materials are relatively inert, allowing operation with both aqueous and hydrocarbon fluid; the authors did not speculate on the capability of the micropump to operate dry (or to pump gases), nor did they mention the particle tolerance of the device. Possible tenfold improvements in pump performance were expected by using larger, single-crystal piezoelectric elements, as well as multiple actuators. This investigation direction seems extremely promising, but unfortunately we are unaware of the current status of the project. An example of highly-integrated micropump was realized using silicon material and electrostatic actuation [40]. In this case, by using multiple stacked wafers, the authors were capable to completely separate the electrically active parts of the system from the fluidic parts, thus avoiding problems related to hydrolysis. Pumping was achieved using a flexible pumping membrane that was attracted electrostatically to a flat parallel electrode. Interestingly, the authors exploited the resonance of the


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Figure 2.19 The integrated plastic micropump manufactured by the company thinXXS. The pump membrane is actuated using a piezoelectric disc, at a relatively low frequency (10–30 Hz). (Used with permission from thinXXS Gmbh.)

valve structures to create reversible pumping action: when the pumping frequency was increased past the valves’ resonance frequency, the out-ofphase operation resulted in reverse. The performance of the valve in forwardpumping was impressive—a maximum 30 kPa back pressure at zero flow and up to 160 µl/min flow rate under no back pressure for a low 5 mW power consumption. The corresponding pumping efficiency was of approximately 0.4%, and the dimensions of the pump were of 7×7×2 mm3 . While this technology seems very promising from an integration perspective, little work seems to be currently underway in this area. A number of commercial micropump models with outstanding performance characteristics are being manufactured by thinXXS Microtechnology [69], a German company that spun off from IMM Mainz. The manufacturing of the valve involves precision injection molded and laser-cut COC plastic for the fabrication of the valve components (housing, valve membrane, and inlet and outlet check valves), and the pumping action is performed by a piezoelectric bimorph disc that is actuated at a relatively low frequency (tens of hertz). The pump achieved Φmax = 9.2 ml/min when actuated at 30 Hz, and a maximum pressure δpmax = 55 kPa in water. This resulted in an impressive maximal pumping power of 2.1 mW, for a power consumption P =250 mW (corresponding to a relatively high pumping energy efficiency η =0.84%). The package size is relatively large (25 mm diameter cylinder, approximately 5 mm in height), which is mostly due to the relatively large piezoelectric

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Figure 2.20 The pump developed by DEBiotech is made in silicon and glass technology, and it is targeted at precise pharmaceutical and medical dosing applications. (Used with permission from DEBiotech.)

disc. The micropump materials limit its operation to fluids compatible with the COC plastic, which excludes most organic solvents and hydrocarbon fluids. Being made of compliant materials, the inlet and outlet valves tolerate relatively large particles (up to 10 µm), which makes the pump interesting for biological and medical applications. The pump can be integrated with a number of other elements on a single chip—leakproof connections and multilevel chips can be made by using laser welding in conjunction with a combination of absorbent and transparent COC material. This micropump is the only commercial autonomous pump solution that is currently available, and the technology is further described in Chapter 5. Another example of micropump is being manufactured by DEBiotech (a Swiss company) in partnership with ST Microelectronics, under the trade name Nanopump. The pump (shown in actual size in Figure 2.20) is targeted towards biomedical drug delivery applications and is based on a design inspired from the early work of van Lintel [38] represented in Figure 2.16. It relies on a piezoelectrically actuated silicon membrane and on two check valves realized in SOI-glass technology [70, 71]. The pump’s application area requires very precise dosage volume per stroke rather than an extremely high pumping pressure or flow rate. In order to optimize this parameter, a double limiter concept was implemented: the stroke of the pump is delimited by two membrane stops, such that a precise volume (160 nl ±5%) is delivered at each stroke despite variations in pressures at the inlet and outlet (within a 100 mbar range), and independent of fluid properties. In order to achieve these goals, the piezoelectric actuator is overdriven and the pump is operated at low frequencies (on the order of 1 Hz). While this does not result in optimal efficiency, it does assure that the membrane performs a full cycle at each actuation. The pump has extremely good longevity: stroke volume remained


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Figure 2.21 The periodic peristaltic sequence: fluid being sucked in from the inlet (left), squeezed between two rollers and the pumping head wall (center) and released at the outlet (right). (Used with permission from ColeParmer.)

constant over more than 2 years when using clean water as the working fluid. The authors do not comment on the particle tolerance of the pump, which is necessarily limited due to the hard materials used in the fabrication of the valves. However, given the drug delivery application that typically employs clean fluids, this is not an issue. By providing fluid contact exclusively with silicon oxide or glass, the pump is chemically inert to the large majority of drugs and chemicals in common use, and is very versatile. The pump element itself is very small (approximately 6 × 10 × 1 mm3 ), but the actuator and electronics require additional volume. While the pump is not available as an off-the-shelf commercial item, it is available for licensing. More details about the technology can be found in Chapter 5. 2.3.3.5 Mechanical Pumping / Peristaltic

Peristaltic pumps employ a mechanism of fluid pushing through a tube or through other elastic medium: a periodic mechanical compression of the tube wall squeezes the fluid out, generating flow. A typical peristaltic pump implementation employs a microtube made from a rubber or silicone material, which is continuously squeezed using rotating cylindrical rollers. The squeezed tube elastically expands after the passage of the roller, sucking fluid in from the inlet. This fluid is then pushed to the outlet by the next roller and the action repeats periodically; a typical configuration involves three rollers encased in a round pumping head cavity, with the tube being compressed between the rotating rollers and the pumping head walls (Figure 2.21). Naturally, the periodic nature of the squeezing mechanism used in peristaltic pumping results in a pulsating flow—at low pumping frequencies the volume of a fluid pulse depends exclusively on the inner diameter of the tube and on the distance between consecutive rollers. The resulting flow rate

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is then given by the product between pulse volume and pumping frequency. Used in this regime, the peristaltic pump can deliver a flow rate accuracy of 3–5%, which is appropriate for some dosing applications. At higher operating frequencies, where the elastic time constant of the tube become comparable with the period of the squeezing action, a frequency dependence of the pulse volume will appear and the flow rate stops being proportional to the pumping frequency. Thermal effects can also affect the accuracy of the pulse volume— at higher pumping frequencies, the mechanical energy dissipation within the pump head results in a gradual increase of the tubing temperature, which can lead to a corresponding change in the elastic properties of the tubing medium. To achieve high accuracy in a peristaltic pump device, it is essential to use precision-extruded tubing because good tubing ID tolerance results in more reproducible flow rates. It is also important to operate the pump at a frequency that allows the tubing to completely recover its shape after the passage of a roller; depending on the tubing material used, a break-in period may be required after tubing change as the shape memory of the tubing material stabilizes. Careful calibration of the pump can help reduce the flow rate uncertainty to below 1%. Planar microfluidic implementations of the peristaltic pumping principle exist. The most widely used technique is realized using PDMS (a flexible silicone elastomer), three-stage pneumatic actuation, and channel occlusion through a flexible integrated membrane [23]. The fabrication technique involves bonding three separately cured PDMS layers: a bottom channel, the separation membrane, and the upper pneumatic channels (Figure 2.22 (a)). Actuation of the channels can be performed using several signal patterns— the pattern used by the original authors was 101, 100, 110, 010, 011, 001. Frequency dependence of the resulting flow rate was linear up to approximately 50 Hz (which corresponded to a flow rate of 1 µl/min). At frequencies between 100 and 300 Hz the flow rate was relatively constant (1.3 µl/min), and at higher frequency it progressively decreased. This effect is due to the relatively slow shape recovery of the peristaltic separation membrane—the lag between pneumatic actuation and complete valve opening is approximately 20 ms (Figure 2.22 (b)). The enoromous advantage of this technique consists of its relatively simple and fast implementation in a laboratory without requiring expensive fabrication equipment: a researcher can easily design and fabricate a complete fluidic circuit involving PDMS pumps and valves in only a couple of days from concept to prototype. Similar technology has been used to implement a two-stage pump using the actuation sequence 00, 10, 11, 01. Surprisingly, the performance of the two-stage pump falls not too far behind that of the three-stage pump [72]. Due to the completely open 11 position in the actuation sequence, however,


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Figure 2.22 A: Soft-lithography peristaltic pumping implementation: a sequence of three pneumatic channels separated from a liquid channel by a thin PDMS membrane. When the channels are actuated, the membrane deforms squeezing the fluid in the liquid channel. B: Corresponding measurements of pressure and valve opening for a 10 Hz pneumatic control signal. (Reproduced with permission from [23].)

such a pump will be subject to some backflow. The efficiency of the pump will therefore strogly depend on the duration of the 11 position, and hence on the pumping frequency. Microfluidic peristaltic pump implementations using harder materials have also been reported. The pioneering work of Smits [39] describes a threestage piezoelectrically actuated microfluidic peristaltic pump using silicon microfabrication technology. The pump employs etched silicon diaphragms with attached piezoelectric disks, which bend under electrical actuation (Figure 2.23). When in rest position, the membrane creates a seal, thus preventing black flow through the device. The membrane pretension and the quality of the seal have a strong effect on the backflow leak rate. The hard nature of the materials (silicon, glass, and silicon oxide) makes good sealing difficult, especially in the presence of particulate matter. While under no back pressure, the pump achieved flow rates of 100 µl/min at 15 Hz actuation frequency; however, the maximum pressure it generated was limited to only 60 mbar. Another implementation in hard materials involved quasi-peristaltic

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Figure 2.23 The peristaltic micropump realized by Smits [39] using silicon and glass as structural materials, and piezoelectric layers for actuation. The threestage actuation sequence is clearly represented. (Reproduced with permission from [7].)

pumping using thermo-pneumatic actuation, with the heater fabricated in silicon, and the peristaltic membranes and flow channel made of machined acrylic plastic material [73] (the pump actually used an asymmetric design with one large volume pumping chamber and two smaller volume valve chambers). The pump, using the same type of three-stage actuation sequence as described above, was capable to achieve pumping rates of 2–3 µl/min under no back pressure conditions. The flow rate was mainly limited by the very slow filling time of the central pumping cavity. 2.3.3.6

Ion-Induced Flow Generation

Several pumping mechanisms based on induced ion motion have been demonstrated. Typically these use an ionic aqueous solution coupled with a combination of electric and magnetic fields to generate ion motion and, hence, flow. The most common types of pumps are electro-osmotic (EO), electrohydrodynamic (EHD), and magnetohydrodynamic (MHD) [7, 8]. 2.3.3.7

Electro-Osmotic Pumping

Electro-osmotic pumps are based on the electric force that acts on the electrical double layer forming at the interface between the channel wall and the ionic fluid. The great advantage of electro-osmotic pumping over all other pumping methods discussed in this chapter consists of the flat velocity profile it can generate, as shown in Chapter 3. Such velocity profiles avoid dispersion


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issues (such as Taylor dispersion) that can significantly degrade resolution in certain analytical applications. Electro-osmotic pumps are also relatively easy to implement: in the simplest form, they involve a capillary connecting two reservoirs with electrodes immersed in them. Application of a voltage difference between the electrodes creates flow between the reservoirs (it is important to note that an electrochemical reaction needs to be possible at the electrodes so that the electro-osmotic flow can be sustained). Electro-osmotic pumps can generate both high flow rates Q and significant pressure differences ∆p. It is important to remember some scaling laws for electro-osmotic pumping [74]: Q ∼ d2 E and ∆p ∼ EL/d2 , where E is the magnitude of the electric fields along the capillary, and L and d are the capillary length and diameter, respectively. We can see, interestingly, that in EO applications the flow velocity (and hence travel time) does not depend on the diameter of the channel; this is a very important point in favor of EO pumping, particularly where very small capillaries are involved. By comparison, in the pressure-driven flow, the maximum flow velocity generated for a fixed pressure difference is proportional to d2 , which has drastic effects on the travel time as d → 0. We also observe that the maximum pressure is inversely proportional to the square of the channel diameter: this observation suggests that reductions in effective channel diameter can lead to significant increase in the maximum pressure capability of the pump. For this reason, EO pumping has been investigated for its potential to generate flow in beadpacked HPLC columns and other porous media [75]. Quite naturally, due to the intimate interface interaction between the fluid and the wall materials, which is responsible for EO pumping, the pumping efficiency will depend strongly on the fabrication materials as well as on the properties of the fluid: pH and ionic strength. Several examples of EO pumps that are capable of driving flow in microfluidic channels exist [7]. Among these, remarkable performance was achieved by pumps using a porous EO medium, such as glass bead-packed systems [75–78], which compensate the reduction in effective fluid channel diameter d by a large cross-sectional area (corresponding to a large number of fluid channels in parallel) to increase Q and obtain a high-pressure, highflow rate system (unfortunately, the dead volume of such a system also increases proportionally). Typical performance achieved by such pumps is ∆p ∼ 100 kPa, Q ∼ 1 − −10 ml/min, when operated at V ∼ 100V. Pumping efficiencies of EO pumps are typically on the order of 1%; it is important to note that the pumping efficiency decreases with increasing ionic concentration as the current flowing through the pump increases as well for a fixed voltage difference V .

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Figure 2.24 Schematic diagram of ACEO pump design using asymmetric electrodes. The net flow direction in the channel is from the narrower electrode to the wider one.

2.3.3.8

AC Electro-Osmotic Pumping

AC electro-osmotic (ACEO) pumps rely on the use of arrays of interdigitated electrodes with alternating potential. The double layer that forms at the electrode surface interacts with the tangent component of the electric field, thus creating flow. If the interdigitated electrodes are chosen to be assymetrical, the flow pattern above the electrodes becomes assymetrical as well, resulting in a net fluid flow (Figure 2.24). ACEO pumps based on assymetric interdigitated electrodes integrated within microfluidic channels have been demonstrated using both straight [79] and circular channels [80]. The net flow could be improved by using interdigitated electrodes patterned on both sides of a channel, in which case a plug flow profile can be achieved [79]. Velocities as high as 500 µm/s have been obtained in ACEO pumping applications. The lifetime of ACEO pumps can be severely limited by electrochemical reactions, which tend to decompose the analytes and erode the electrodes over time. In one study, the lifetime of 70-nm-thick gold electrodes was as low as 90 minutes; ITO, on the other hand, appeared to be less affected by electrolysis [80]. Electrode erosion is accentuated by driving the pump with high voltages. 2.3.3.9

Electro- and Magnetohydrodynamic Pumping

EHD and MHD pumping, while used in certain applications, have not seen nearly as much development as the other pumping methods described in this section. Several reasons may be responsible for this, most importantly the relatively poor performance of such pumps compared to alternative techniques. For example, the maximum pressures that can be generated are


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usually in the neighborhood of ∼100 Pa, as compared to the much higher pressures that can be generated by EO pumps or any of the reciprocating and peristaltic pumps we described earlier. The basic working principles of the EHD pumps rely on the induction of charges at a material interface, such as the interface of two dielectric fluids. By imposing a traveling wave electric field pattern using an array of electrodes, the charges are pulled along the interface, creating a net flow of fluid in the same direction. Other principles involve ion injection at the electrodes, which are then carried by the electric field (by electrophoresis), thus generating bulk flow. Finally, MHD pumps relay on the Lorenz force on traveling ions between electrodes disposed transversely across a channel and an external perpendicular magnetic field. The Lorenz force then acts along the channel of interest, and the moving ions generate bulk flow in this direction. It is important to note that, unlike EO pumps, both EHD and MHD pumps act on the bulk of the fluid and will generate flow profiles similar to pressure-driven flow, and therefore Taylor dispersion will be an issue. Another important point to make regarding ion-induced flows is that significant Joule heating of the fluid may be generated by the ionic currents. While in very narrow geometries this may not be an issue, as the surface-to-volume ratio of the walls is relatively important, thus allowing efficient heat sinking, other types of geometries may generate a nonnegligeable heating of the working fluid. Since relatively few examples of efficient EHD/MHD pumps exist (particularly with applications for creating highly integrated devices), we will not discuss specific implementations of these pumping principles in detail here. Interested readers are referred to the review articles mentioned at the beginning of Section 2.3. 2.3.4 Single-Phase Mixing Mixing in microfluidic devices is generally difficult. The low Reynolds numbers typically appearing in microfluidics imply that turbulence (the mechanism responsible for most mixing phenomena at the centimeter scale and above) is unlikely to develop in such devices. The laminar nature of the flow implies that diffusion will be the main mechanism for performing good mixing. Since diffusion is relatively slow, specific mixing geometries have been developed that can shorten the diffusion length and, consequently, the species’ mixing time. Typical approaches followed in microfluidics involve either generation of chaotic interfaces or of laminated fluidic structures between the fluids of interest by convective means—the diffusion length is therefore decreased significantly, which leads to an important reduction in mixing time. We will discuss several of these approaches next; however,

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Figure 2.25 Different capabilities of flow patterning. (a) Two parallel (unreactive) dye streams, showing the relatively low mixing even at several millimeters downstream from the junction point. (b) Etching pattern in a gold layer preexisting on the substrate. (c) Deposition of a silver microwire at the interface between components of an electroless silver plating solution. (Reproduced with permission from [86].)

for more detail, the reader is referred to several review articles [81–85], as well as to general books on microfluidics that treat at length the subject of micromixers [1, 5]. Heavy mixing in microfluidic devices is not always desired—sometimes, what is desired may be parallel streams of two different liquids that only mix in a narrow interfacial region. The laminar nature of flow in microfluidic devices naturally provides this capability. Flow patterning is an application where this feature may be useful: two chemical species are injected as parallel streams in a microchannel, and they only mix (and react) along the interface, thus providing a potential route for microfabrication with accurate positional and dimensional control [86]. The reaction product may deposit at the bottom of the microchannel, which provides a micron-scale fabrication route for different structures, most notably metallic lines; alternatively, the reaction product may itself react with the chip materials, thus etching very locally to pattern or structure the walls of the channel (Figure 2.25). While this fabrication route has potential in very specific application, its capability to create complex patterns is heavily limited by the liner topology of the microchannels, and does not have immediate applications to fabricating labon-a-chip devices. It will therefore not be explored further.


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In the majority of the applications, mixing needs to be enhanced. This reduces, for example, the time required for a chemical reaction or analysis to be performed, with important effects related to the throughput of the device. In-line analysis systems are particularly sensitive to the mixing issues: as the fluid passes continuously through the device, only a limited time (given by the device geometry and the flow rate used) is available for performing mixing and measurements on the fluid stream. Several mixing strategies can be employed, and these are roughly divided into passive or active techniques. Passive techniques rely on interactions between imposed flow and certain geometrical features of the channel. Active techniques, on the other hand, employ actuators (mechanical, electrical, thermopneumatic) that agitate the fluids in a direction normal to the mixing plane, thus creating a wavy interface with larger area and consequently enhanced diffusive transfer. 2.3.4.1 Passive Single-Phase Mixing

Irreversible processes, such as molecular diffusion and chaotic flow (which appears only in the turbulent flow regime), are required in order to produce a fully homogenous output stream from several distinct input streams. As the Reynolds number in most microscale applications is low (significantly below the turbulent transition), turbulence rarely plays a role. The vast majority of micromixers therefore need to rely on diffusion as the irreversible mixing step. Without diffusion, mixing would not be achieved at very low Reynolds numbers. Indeed, due to the reversibility of the flow equations in this case, reversing the velocity of the fluids would imply that from the final state, one would be able to reconstruct the initial, fully separated state. This is impossible if the final state is fully mixed (i.e., homogenous). As we will learn in Chapter 3, the diffusion time across a distance l is given by l2 tD ≈ (2.1) D where D is the diffusion coefficient of the species of interest. This time can be considerable, reaching several minutes for mixing a small molecular weight aqueous solution across a 500 µm channel (D ≈ 10−9 m3 /s, and hence tD ≈ 250 s). Considering that the stream velocity at the interface between the two fluids is given by v, the width of the mixed p region lm at distance d downstream from a T-junction is given by lm ≈ Dd/v. In turn, the channel length d required to achieve complete mixing of two adjacent substreams in a square section channel of a cross-sectional dimension w is given by d ≈ vw2 /D ≈ Φ/D where v Φ is the volumetric flow rate. For a typical flow rate of 10 µ l/min, this corresponds to a minimum channel length d ≈17

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cm, whereas for mixing solutions of larger proteins (D ≈ 10−10 m3 /s), the required channel length can be of several meters. From (2.1) it becomes apparent that the total time required for full mixing can be reduced considerably by reducing l: reduction of l by a factor α will therefore reduce the diffusion time by a factor α2 . This effect is used in many passive mixer devices: the flow is manipulated in such a way as to create thin laminated layers from the input stream, which then can benefit from the shorter diffusion distance to achieve faster mixing. Alternatively, transverse flow across the channel is generated by different means to create a wavy interface between the streams, with the same overall effect of enhancing the contact surface and reducing the total diffusion length. The parallel lamination of multiple streams as an effective mixing strategy in microsystems has been studied for over 15 years, for microchemical reactor applications. Lamination of two streams into n parallel substreams will reduce the mixing time by n2 , as we have seen above. Different types of lamination designs have been proposed. In one implementation, the two streams to be mixed are connected to a network of 15 dead-end interdigitated channels [87]. The resulting laminated flow leaves the device in a direction perpendicular to the plane of the original streams, and the theoretical mixing diffusion time is reduced by 225 times. The device was fabricated using metallic materials and a combination of LIGA and micro-EDM techniques for microfabrication (Figure 2.26). While the specific device described here was designed as a stand-alone unit and manufactured using technologies with little integration potential, the concept could be used successfully in other implementations to achieve rapid mixing in an integrated device. Another technique for achieving lamination involves successive doubling of the number of layers by using sequential lamination procedures (also called a baker’s transformation by analogy with the techniques used in making a homogenous dough by stretching and folding—the Reynolds number in the baker’s case is also very low). An example, shown in Figure 2.27, describes a simple procedure for achieving a laminated fluid layer structure containing twice the number of initial layers. The diffusion time required for full mixing, in this case, divides by a factor of 4 at each such lamination step. By using n such lamination steps sequentially, the diffusion time required for full mixing will be reduced exponentially by a factor of 4n . While in theory it might be possible to split each channel in more than two subchannels, and thus triple, or quadruple, the number of layers at each lamination step, in practice this is very difficult to achieve using standard planar microfabrication technology. If faster mixing is required, it is simpler to cascade additional layer-doubling steps. Different implementations of such lamination techniques have been studied and reported in the literature, most often using silicon/glass micromachining


a

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b

Figure 2.26 Top: A schematic representation of the fluid lamination principle: by doubling the number of parallel fluid layers, the diffusion length required for full mixing (represented by vertical arrows) is halved (Panels A, B, C). The multiple parallel layers of fluid finally mix by diffusive processes (Panel D). Bottom: Micromixer device relying on the principle of multiple stream lamination to reduce the diffusion length. The microfluidic part of the device was realized in metal (nickel and silver) using the LIGA fabrication technology, and in other metals using micro-EDM techniques. Panel a: Complete device, with fluidic connections, housing, and mixer units (left: nickel device, right: silver device). Panel b: detail of the mixer unit, which creates a wavy laminated stream from the two input streams in a direction perpendicular to the device plane. (Reproduced from [87], with permission from Elsevier.)

[88–90] or Mylar sheet lamination [91]. Since the channel splitting and reuniting necessary for this technique to work requires careful alignment of multiple layers over the full area of a device, it is difficult to implement it using soft-lithographic techniques (we are unaware of any such implementations). Chaotic advection can be induced by certain wall geometries in microchannels. This phenomenon has been used for generating mixing by using asymmetric herringbone-like protrusions on the walls of a microchannel [92]. The induced laminar chaos generates a wavy spatial interface between the fluids to be mixed. The surface of the interface and the width of the resulting interdigitated layers exponentially increases and respectively decreases with downstream distance (or, equivalently, with time). From this point of view, this mixing technique is similarly effective to the sequential

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a

b

Figure 2.27 The sequential lamination mixing technique consists of multiple “layerdoubling” steps. Such a step consists of splitting a channel containing two adjacent streams (Panel a1) along a line perpendicular to the interface (Panel a2), then expanding the resulting half-channels in the vertical direction to obtain two parallel channels (Panel a3), and finally joining the two channels into a single channel to obtain four adjacent substreams (Panel a4). At each such step, the number of fluid interfaces doubles, and the diffusion time required for full mixing reduces by a factor of 4. An implementation of such a device in silicon using DRIE micromachining is shown in Panel b. (Reproduced from [90], with permission from Elsevier.)

lamination method described above. While the wall-induced chaotic advection technique does not guarantee a uniform mixing across the full cross-section of the microchannel [81], it has an important advantage over the sequential lamination technique in that it can be implemented using simpler fabrication technology (including soft lithography). Similar chaotic advection can be generated using heterogenous surface charges for electrokinetic flows [93]. Flow focusing is another technique used to enhance mixing in microfluidic channels. By using tapered channel geometries with progressively reduced width in the direction perpendicular to the mixing interface, the total mixing time can be reduced proportionally at the expense of increased channel hydrodynamic resistance. Alternatively, hydrodynamic focusing can be used by sandwiching the solute stream between two identical solvent streams (Figure 2.28). While this technique is usually applicable in cases


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Figure 2.28 Hydrodynamic focusing can be used to generate mixing on time scales as short as a few microseconds. In the figure, the middle stream was fluorescently dyed using fluoroscein. (Reproduced from [94].)

where relatively low concentration solutions are required, it does not lead to increased hydrodynamic resistance and can be a very effective mixing method; mixing times down to a few microseconds have been reported [94]. The methods presented in the above section can be combined in a single device to obtain even faster (or more efficient) mixing. For example, mixing between streams of different viscosities may be achieved by hydrodynamic focusing initially (which assures that the mixing interface lies in the fastflowing region from the middle of the channel rather than near a wall) followed by successive steps of flow folding and wall-generated chaotic advection to obtain a fully mixed, homogenous output stream. 2.3.4.2 Active Single-Phase Mixing

By contrast with passive mixing devices, active devices use some form of external actuation to generate either laminar chaos in the microchannel, or a wavy, interdigitated interface between the streams by creating periodic transverse flow. While passive mixing has received quite a bit of attention and has been studied in detail, active mixing, by comparison, has seen less scientific activity. Implementing active micromixers is significantly more

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expensive than any of the passive techniques described in the previous section, and their use makes economic sense only when the added value of using an active component outweighs the costs. This may be the case in a highly integrated system, where other components that use similar technology need to be included, the cost of adding the active mixer block being only marginal. One technique of creating a periodic interdigitated interface between two streams is to contact the two streams in a T-junction, and to alternately inject fluid from each stream [95] (Figure 2.29). This creates effective plugs of the two fluids, which stretch and mix due to Taylor dispersion and longitudinal diffusion. The technique is best implemented using local flow control, either by using an external pressure source and integrated valves driven out of phase on the input streams, or simply by using integrated micropumps where, again, the actuators are driven out of phase. Completely external flow control is also possible; however, effects due to compliance issues and dead volumes may affect the timing of the fluid injection and thus reduce the mixing efficiency. A similar technique uses a cross junction, whereas side-by-side fluid streams are injected from one channel, and two side channels are pulsed to periodically move the fluidic interface in the transverse direction. The stream exiting from the fourth channel in the cross shows a wavy interface between the two streams. By stretching and folding fluid in the main and side channels, chaotic mixing can be achieved. This type of mixing process, based on periodic displacement of the fluidic interface due to external flow, has been studied theoretically by several groups [96, 97], and it has been implemented using different actuation mechanisms: magnetohydrodynamic forces [98], pneumatically and thermopneumatically generated flow [99–101], and pressure pumping [102, 103]. Different fabrication technologies were used (PDMS soft lithography, silicon-glass micromachining)—the mixing technique can be integrated with any pumping or valving technique that uses similar actuation mechanisms. Electrokinetic instabilities in electro-osmotic flow driven by AC electrical fields have been used to generate mixing [104]. An alternating electric field has been applied tangent to the interface between two substreams generated using a T-junction. One of the fluids was fluorescently dyed, simplifying quantitative optical investigation. When the frequency of the electric signal was on the order of a few hertz, an instability developed that generated chaotic flow within the microchannel, thus mixing the two substreams. It is interesting to note that the time scale for achieving a reasonable degree of mixing was on the order of several seconds, thus much longer than the period of the AC excitation.


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a

b Figure 2.29 Alternating flow from two out-of-phase streams in a T-junction geometry creates a wavy interface between the streams, enhancing mixing. (a): Actual device geometry. (b): Results of numerical simulations showing improved mixing downstream from the T-junction. (Reproduced with permission from [95]. Copyright 2004 American Chemical Society.)

2.4 Droplet-Based Fluid Manipulation As opposed to the continuous flow microfluidic devices described above, droplet-based microfluidics relies on the manipulation of individual droplets of a liquid that are immersed in an immiscible fluid (which could be either a liquid or a gas). The droplet diameter in a liquid-liquid system can be made larger than the channel diameter (in which case the droplet becomes a fluid plug touching the walls of the channel), of similar size (with very limited contact with the walls), or even smaller (no contact with the walls); the majority of droplet based microfluidics involve relatively small droplets of the same order of magnitude as the channel diameter. The two-phase nature of the flow in droplet microfluidics represents a significant complication relative to single-phase flow: the effective viscosity of a two-phase medium depends

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Figure 2.30 Example of fluid boundary being disturbed by the electro-osmotic flow field generated by an alternating electric field. A hydrodynamic instability develops that is responsible for the observed chaotic mixing. The AC frequency of the electric field was 10 Hz. (Reproduced from [104]. Copyright 2001 American Chemical Society.)

on a large number of parameters such as the viscosities of the two separate phases, size of the droplets relative to the channel diameter, the ratio of droplet to continuous phase volumes, the type of surfactants used, and so forth. In addition, surface and wetting effects such as Marangoni stresses or contact line dynamics can play an important role in multiple phase systems. Depending on the complexity of the design, such issues may be just irrelevant nuisances, or they may represent significant obstacles that require a complete redesign of the system. The interest of droplet-based microfluidics is in the use of individual droplets to perform certain functions in a completely isolated environment, thus avoiding Taylor dispersion issues common to single-phase microfluidics: a droplet could be an individual chemical reactor, millions of droplets representing as many experiments; alternatively, a droplet can contain biological material such as DNA, and serve as a vehicle and reaction vessel for certain genetic manipulations such as PCR (polymerase chain reaction), a laboratory technique commonly used to amplify nucleic acid sequences. Alternatively, microdroplets generated in microfluidic systems can be used as vesicles for drug encapsulation, in personal care products, or as precursors for more complex materials such as colloidal crystals.


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A number of techniques specific to droplet-based microfluidics have been developed, and are described in Sections 2.4.1 to 2.4.3. Typical individual operations are droplet formation, mixing, splitting, coalescence, switching, trapping, and separation. In addition to the publications cited in the text, a number of excellent review articles covering droplet-based microfluidics exist [105–108]. The reader is encouraged to consult these for additional information about the techniques described in the following sections. In addition, Chapter 5 describes a commercial application of droplet-based microfluidic technology. 2.4.1 Droplet and Emulsion Generators Droplet generation implies a transformation of one continuous stream of liquid into individual droplets. In microfluidics, where inertial effects are relatively unimportant compared to surface tension, this transformation is mediated by a capillary instability—for droplets to form, they must reach a state where splitting from the continuous streams is energetically favorable. Several methods have been developed to generate microdroplets in a microfluidic system. The simplest configuration involves a T-junction, whereby a dispersed phase is injected, via a side channel, into a stream of the continuous phase. Droplets obtained by this technique typically result in a fluid plug; to obtain droplets smaller than the channel, one can reduce channel sizes in the T-junction region and then enlarge the channel downstream. The droplet formation process repeats periodically as more continuous and dispersed fluid is injected into the T-junction. The size and monodispersity of the resulting droplets depend on the relative pressures (or flow rates) of the two fluid streams (Figure 2.31), on their viscosities, and on the type and concentration of surfactant used [109, 110]. Several mechanisms exist that can explain the droplet breakup at a Tjunction. The flow of continuous phase advects the fluid downstream as an incipient drop still connected to the dispersed fluid stream by a thinning fluid thread. At some point the thread becomes energetically too costly and undergoes a capillary instability, thus freeing the droplet. The process is similar to shear-rupturing techniques for generating emulsions in bulk systems [111]. The above mechanism correctly describes the droplet breakup mechanism for moderate capillary numbers. In the low capillary number regime, however, the interfacial forces dominate the shear stress, and the breakup occurs due to a pressure difference that develops across the droplet as it forms. In this regime, the droplet size is determined almost exclusively by the flow rates of the two streams and not by the fluid properties [112].

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Figure 2.31 Formation of droplets at a microfluidic T-junction, showing the dependence of the droplet size on the relative flow rate of the dispersed phase (water in this case). (Reproduced from [110]. Copyright 2003 American Chemical Society.)

Other configurations commonly used in microfluidic devices involve flow-focusing geometries. In a flow-focusing approach, the dispersed stream is surrounded on all sides by the continuous phase stream, and both are injected through a nozzle. The dispersed flow forms a filament and then breaks into droplets (Figure 2.32). The breakup is mediated by the RayleighPlateau instability: a thin fluid thread becomes unstable to small perturbations (which grow and eventually sequence the thread into drops) when its aspect ratio increases past approximately 3.15. The size and frequency of the droplet generation depend on several parameters, such as the device geometry, the relative flow rates of the phases, the viscosities, and the surfactant concentrations. An interesting aspect of flow-focusing is that it can readily generate droplet sizes that are smaller than the smallest system dimension [113] by using an entirely planar geometry requiring a single photolithographic step. This feature makes flow-focusing an easy-to-integrate brick in almost any fabrication technology. The typical diameter of the droplets commonly generated by flow-focusing devices is on the order a few to a few tens of micrometers in diameter, with a polydispersity on the order of a few percent; the drop production frequency can range from tens of hertz to


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Figure 2.32 Left: Formation of droplets via flow-focusing in a planar microfluidic geometry. The continuous phase (oil) is injected form the top and bottom channels at flow rate Qo , whereas the dispersed phase (water) is injected through the middle channel at flow rate Qi . Both are forced through the nozzle, which produces different droplet configurations downstream, depending on the relative flow rates Qo and Qi . Right: A similar flow-focusing geometry in three dimensions allows the addition of a third fluid stream, resulting in a controlled double emulsion generation. (Reproduced with permission from [113] (Copyright 2003, American Institute of Physics) and [114], respectively.)

more than 10 kHz [106]. Flow focusing can also be used in three-dimensional microfluidic devices, to create complex emulsions such as double emulsions involving droplets of one fluid enclosed in droplets of a second fluid, which are in turn enclosed in larger droplets of the first fluid. To this end, the simple emulsion generated by a first flow-focusing device can be injected as the dispersed phase in a second flow-focusing device, or alternatively, multiple fluid streams can be injected into the same flow-focusing device. The resulting double emulsion droplets can have a controlled number of single-emulsion droplets trapped inside of them [114]. More recently, planar devices have been fabricated using soft lithography techniques and wetting control. These devices allow a cascade of multiple flow-focusing junctions to create simple, double, triple, quadruple, and even quintuple emulsions [115]. In order to further control the size of generated droplets, an electric field can be applied to the flow-focusing junction in the direction of droplet ejection [116]. Figure 2.33 shows the effect this has on the resulting droplet diameter—as the electric field is increased, a Taylor cone is generated at the tip of the dispersed phase stream, with the size of the generated droplets strongly

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Figure 2.33 Droplet formation at a flow-focusing junction (Qi /Qo = 1/17.5) as a function of applied voltage, shown in the lower right corners (the corresponding electric field is parallel to the vertical axis of the images). The size of the droplets decreases from a few micrometers when no field is applied to submicron diameters at the maximum electric field. (Reproduced from [116]. Copyright 2007, American Institute of Physics.)

depending on the imposed electric field. Submicron droplets can be generated using this technique. Other droplet-generation techniques that have been employed in microfluidics involve the topological transition of a droplet from two to three dimensions, as is the case in microchannel emulsification (MC) [117]. A droplet is generated using a T-junction geometry on a shallow terrace bounded by two plane walls. The droplet develops a planar “pancake” shape, with the thickness of the pancake given by the spacing between the top and bottom terrace walls. At a certain distance from the T-junction, the terrace disappears, so that the droplets undergo a 2-D to 3-D transition. Since the 3D spherical droplet shape is much more favorable energetically, the capillary instability at the 2-D to 3-D interface is abrupt and droplet formation is very reproducible (resulting in good monodispersity). As compared to microfluidic flow-focusing techniques, which are limited by the planar device topology to just one or a few flow-focusing junctions per device, the MC technique allows hundreds of channels to produce droplets in parallel, with much improved throughput (Figure 2.34, left). To further increase throughput, the channels and corresponding terraces can be manufactured perpendicularly to the substrate [118], thus allowing fabrication from a two-dimensional array of channels (Figure 2.34, right). In both cases, the resulting polydispersity is on the order of 2%. An original and relatively recent technique to generate and at the same time to store the resulting droplets involves the use of capillary effects, in the form of chip regions that allow the flow of the (wetting) continuous phase, but block the flow of the (nonwetting) dispersed phase due to the formation of a capillary meniscus. The droplet generation functionality is most easily


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Figure 2.34 Left: Creation of droplets at a 2-D to 3-D transition in microchannel emulsification. Droplet generation is performed in parallel through a linear array of channels. Right: A similar geometry, where the channels and terraces for droplet generation are oriented perpendicular to the substrate, forming a planar array. (Reproduced from [117] and [118], respectively. Copyright 2001, 2005 American Chemical Society.)

achieved by using small diameter restrictions at the end of larger diameter channels coupled with bypass channels (Figure 2.35). This technology allows the generation and storage of a large number of droplets (in the form of plugs) on-chip [119], without the need for flow-focusing geometries. The speed at which the droplets are formed is limited, however, by the fluid pressures (which must be low, or else the droplets would overcome the capillary pressure barrier and escape the traps through the small capillaries), and by the total fluidic resistance of the system, which is proportional to the number of drops that need to be generated. Several electric-field-assisted droplet-generating techniques have been developed, based either on dielectrophoretic (DEP) or on electrowetting on dielectric (EWOD) techniques. In both cases, the droplets are generated from a reservoir (which could be a macroscopic drop) in contact with either silicon oil or with air. Both techniques are relatively slow compared to any of the previously described techniques: droplet generation can reach a frequency of tens of hertz at most. In DEP droplet formation [120], elongated “finger” electrodes can be placed in the immediate vicinity of the reservoir drop. When a voltage is applied, a narrow stable fluid rivulet is formed above the electrodes. As soon as the electric field is turned off, the rivulet becomes unstable, and breaks into individual droplets because of the Rayleigh-Plateau instability. If the topography of the finger electrode is patterned with a periodic wavelength similar to that of the fastest growing capillary instability (which is roughly

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Figure 2.35 An example of microfluidic device using capillary effects to create and store droplets in a microfluidic device. As the plug of dyed water advances, it enters the regions terminated in restrictions but cannot penetrate through the restriction itself due to the formation of a capillary meniscus. A droplet is therefore formed at each such region, whose volume can be tailored by the length of the channel ending at the restriction. Drops can then be pumped out of the device by reverse flow. (Reproduced from [119]. Reproduced by permission of The Royal Society of Chemistry.)

Figure 2.36 Left: DEP droplet generation. The rivulet that forms along the linear electrodes when activated breaks off into small droplets as soon as the voltage is turned off. The periodic patterning of the electrodes with bumps at the wavelength of the fastest growing instability results in equally spaced droplets of uniform volume. Right: EWOD droplet generation in a water-in-oil system. (Reproduced from [120] with permission from Elsevier and from [121] by permission of The Royal Society of Chemistry, respectively.)

given by 3.15 times the distance between the electrodes, as implied by the Rayleigh-Plateau theory), then the resulting droplet size and spacing are very reproducible (Figure 2.36). By contrast with the DEP technique, which is due to body forces caused


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by electric field nonuniformity, EWOD relies on the modification of the surface energy of the liquid-solid interface, and implicitly of the respective contact angle (a phenomenon called electrocapillarity). In EWOD techniques, the most common configuration consists of conductive electrodes patterned below and above the fluidic channel and covered using a dielectric layer; alternatively, a single plane of electrodes could be used. The mechanism of droplet formation is in fact very similar to the DEP technique described above: several electrodes are activated in a row, starting from the fluid reservoir (this causes a fluid finger to advance and cover the surface of the electrodes completely). Then all the electrodes except the farthermost are switched off. The fluid finger will then break off and retreat to the reservoir, leaving a droplet attached to the active electrodes. The technique has been initially demonstrated using silicon oil as the surrounding medium [121], and more recently has been accomplished in air [122]. Some controversy exists on whether EWOD is really based on electrowetting as the main driving force, or the electrocapillary effect only reduces droplet pinning (or contact angle hysteresis), while the actual driving force is of DEP origin [123]. In any case, a distinction can be unambiguously made: in EWOD, the triple contact line is essential for operation, whereas in DEP systems, no contact line is necessary. 2.4.2 Droplet-Based Mixing Droplets are very interesting to the chemical and biological community as miniaturized reactors: a mixture of chemicals or of biological agents can be encapsulated within a droplet and the allowed to evolve. Due to the relatively small droplet size, diffusion times are reduced and thus reaction kinetics can benefit from a boost. Diffusion alone, however, is too slow for many microfluidic applications—in many applications a droplet is generated every millisecond, so the total mixing time needs to be of the same order. Active mixing needs to be performed in this case to homogenize the mixture of reactive compounds. Droplet-specific mixing mechanisms need therefore to be developed. Two main directions have been investigated to achieve droplet mixing. One method relies on chaotic advection driven by asymmetric lateral wall protrusions [124], or by sharp turns in a serpentine channel [125]. In both methods, one uses the internal circulation that is induced within the droplet due to the boundary conditions. In a straight channel, however, the droplet possesses two planes of symmetry, and therefore the droplets are localized within each quarter. In this case, induced circulation is only effective in mixing the fluid within each quarter (diffusion is required for reactant molecules to hop across a boundary, leading to complete mixing). If any kind of

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asymmetry is introduced in the system, however, the shape of the vortices adjusts accordingly, causing some fluid to cross from the smaller vortices to the larger ones. This explains the mixing observed in the systems shown in Figure 2.37 (top panels). Mixing times as low as 3 ms were measured, which are short enough to enable measurements of chemical reaction kynetics. It is interesting to note that the structure of the vortices, even in a straight channel, turns out to be extremely complex, as the bottom panels of Figure 2.37 show. In certain situations the presence of the droplet can create enough drag that the continuous fluid will bypass the droplet and provide corner flow. The middle panel of Figure 2.37 only represents a simplified description of the phenomenology. The second mixing method relies on the use of electric fields to generate back-and-forth droplet motion, and thus induce mixing [127]. The device is based on dielectrophoretic (DEP) actuation, and involves electrodes patterned on the top and bottom walls of a Hele-Shaw type cell. Water droplets surrounded by silicone oil are moved between electrodes using electric fields (the water droplets are attracted to the high-field regions due to the high dielectric constant contrast between water and oil). The time to complete mixing depends on the number and spatial distribution of electrodes as well as on the motion pattern. As expected, the more complex the motion, the faster the mixing time. Mixing can be achieved in as little as 3 seconds if a pattern of eight electrodes is used and the droplet motion follows a trace containing both clockwise and counterclockwise loops (Figure 2.38). 2.4.3 Droplet Manipulations As we have seen in the previous section, electrical fields can be used to manipulate droplets—the DEP force acts in regions of electric field gradient, and pulls materials with a higher dielectric constant than the surrounding medium towards the high electric field regions (similarly, the electric field can be used to repel materials with a lower dielectric constant than the medium). EWOD techniques, on the other hand, rely on the temporary modification of liquid-solid surface energies under the application of an electric field, thus resulting in a change in the contact angle and a consequent change in the capillary force. By using these types of actuation, it is possible to trap droplets, to move them around between different electrodes (as in the mixing example above), or to deflect their trajectory from one possible outlet channel to another (e.g., to sort them). In one example, particles are deviated from one input stream to one of five possible output streams using DEP forces [128] (Figure 2.39, top), using multilayer SU-8 photolithography in conjunction with metal deposition and


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Figure 2.37 Top left: Asymmetric channel wall protrusions used to induce flow folding in microdroplets, leading to mixing enhancement. Top right: Similar enhancement in mixing due to bends in serpentine channel. Middle: Schematic representation of flow lines within a microfluidic droplet, showing the asymmetric circulation near bends, and the symmetric circulation pattern in a straight channel. Bottom left: Experimental observation of internal droplet circulation, showing strong counterrotating vortices due to the corner flow. Bottom right: Cross-section at a quarter of the height, showing the structure of the counter-rotating vortex. (Reproduced from [124] (by permission of The Royal Society of Chemistry), [125] (by permission. Copyright Wiley-VCH Verlag GmbH & Co. KGaA.), and [126] (by permission of The Royal Society of Chemistry) respectively.)

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Figure 2.38 Mixing using an electrowetting device: droplet motion between the electrodes follows a figure-of-8 pattern containing both clockwise and counterclockwise loops. This assures rapid mixing of the droplet content, as the successive frames shown above demonstrate. Only the clockwise loop is shown. (Reproduced from [127] by permission of The Royal Society of Chemistry.)

electroplating to obtain vertical electrodes on the side of the channel (the vertical positioning of the electrodes was critical in achieving the complex sorting functionality). In a different example, planar electrodes positioned beneath a PDMS channel were used to deviate one droplet from a stream to a given side channel [129] (Figure 2.39, bottom). This type of actuation is very convenient, as the electrical force acts instantaneously, as compared to other types of actuation used in microfluidics. This allows, for example, for very rapid sorting—a single droplet can be extracted from a stream of droplets generated at kilohertz frequency. Droplet coalescence or splitting can be achieved using a number of different techniques. Oppositely charged droplets that are synchronously generated using two electric-field-enhanced flow-focusing junctions have been shown to coalesce reproducibly even in the presence of surfactants [130]. By contrast, droplets that are not charged and not synchronized do not coalesce, but evolve side by side (Figure 2.40 bottom left). Coalescence can be enhanced even between uncharged drops by the application of a local electric field (a phenomenon called electrocoalescence [131]; Figure 2.40, top) or by heating the continuous phase fluid between the drops, which reduces its viscosity and thus enhances drainage.


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Figure 2.39 DEP switching of particles and droplets from a continuous flow stream. Top five panels: Using vertical electrodes integrated in the sidewalls of a channel, it is possible to reproducibly switch the outlet channel of a particle using an electrical signal. (a–c): Electrodes patterned underneath the channel can be used to switch droplets between the two outlet channels by activating the respective electrodes. Sorting can be performed rapidly on streams of droplets generated at kilohertz frequencies. (Reproduced from [128] (by permission of The Royal Society of Chemistry), and from [129] (by permission. Copyright 2006, American Institute of Physics), respectively.)

Splitting can be achieved passively at a T-junction, a situation that has received significant theoretical attention [132, 133]. A droplet reaching a junction will split into two daughter droplets if the conditions are right; the volume of the resulting droplets depends on the relative flow rates in each of the outlet channels [134]. If an electric field is applied to the junction at the moment of the splitting, and the droplet fluid contains ionic species, the initial drop becomes polarized and splits into two oppositely charged droplets. This technique can be used for recharging neutral drops (Figure 2.40, bottom right), which can then be manipulated using electrostatic actuation [130].

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Figure 2.40 Top: Electrocoalescence of neighboring droplets using an electric field pulse. Middle: synchronous generation and coalescence of droplets using two electrically coupled flow-focusing junctions (first panel: scematic of the system; second panel: system behavior in absence of the electric field; third panel: system behavior in presence of the field). Bottom: the generation of charged droplets during splitting at a T-junction in an electric field. (Reproduced from [131] (by permission. Copyright 2006, American Institute of Physics) and [130] (Copyright Wiley-VCH Verlag GmbH & Co. KGaA. Reproduced with permission.), respectively.)

Droplet splitting and coalescence can also be achieved using EWOD technology. The activation of the electrodes on two opposite sides of a droplet can force a droplet to break into two equal-sized daughter droplets. Similarly, activating the electrode between two separate droplets can force them to coalesce [121,122]. Such droplet manipulation steps can be performed repeatedly and reproducibly (Figure 2.41). Active droplet manipulation is also possible using the thermocapillary effect—by heating one side of a droplet, the surface tension is reduced, and droplet motion can be initiated. While the throughput and the practical implementations of this technique are limited due to the slow thermal response of most substrates and liquids and to the inherent droplet heating that occurs, it


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Figure 2.41 Successive droplet splitting (a–c) and coalescence (c–e) operations realized using EWOD droplet manipulations. The state of the electrodes at each step is marked (on, off). The total process time is on the order of 1 second. (Reproduced with permission from [122].)

can be used in a number of applications. The technique has been demonstrated on silicon substrates using complex arrays of integrated heaters to generate the desired temperature gradients [135, 136]. 2.4.3.1 Phase Separation

We have seen earlier several methods for creating droplets and emulsions in microfluidic systems: many of the successful techniques are based on surface tension and capillary instabilities, and controlled emulsions can be generated this way in a microfluidic system. The reverse process, of transforming an emulsion (or more generally, a two-phase stream) into its single-phase components (demulsification), turns out not to be a trivial task; many industries (such as the oil production sector) are routinely confronted with the problem of stable emulsions. Due to the small scales involved in microfluidic technology, it appears as an attractive candidate for phase separation: at dimensions attainable in microfluidic devices surface effects like capillarity become dominant over gravitational and inertial forces, offering interesting routes for fluid

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manipulations that cannot be implemented in larger systems. As an example, gas-liquid separation in a microfluidic device has been reported [137], gas bubbles being separated from a liquid stream by using capillary forces. Capillary liquid-liquid extraction in microfluidic geometries has also been reported [138] in a system for continuous liquid extraction of several organic and fluorinated compounds from a water stream, utilizing an integrated fluoropolymer membrane. Liquid-liquid separation based on capillarity has been demonstrated in a polytetrafluoroethylene (PTFE) device [139]. The physics of capillary separation and droplet breakup in side channels under several geometries realized in polydimethylsiloxane (PDMS) has been investigated [134]. More recently, single-phase sampling and subsequent optical absorption measurements on the separated stream have been performed in a microfabricated system using a PTFE membrane interface [140].

2.5 Detection and Measurement Techniques So far we have learnt how fluids can be manipulated on a microfluidic chip by performing pumping, mixing, valving, heating and cooling, droplet generation, splitting, and coalescence operations. The techniques we described may be sufficient by themselves to perform a certain number of operations, such as biological and chemical sample preparation and conditioning, or fabrication of complex microstructured materials, such as emulsions or colloidal assemblies [141]). Most microfluidic applications, however, require some form of measurement to be performed on the microfluidic chip in addition to simple fluid manipulation. In the following sections we will review the main directions investigated for performing sensing and measurement actions in microfluidic devices. We identify three main classes of measurements that can be performed in a microfluidic device: optical measurements (which can involve fluorescent detection, UV-VIS-NIR absorption spectroscopy, light scattering, refractive index measurements, and so forth), chemical measurements (separation-based measurements and electrochemical cells), and physical measurements (involving the physical characteristics of the fluid that are not of optical origin—viscosity, density, flow rate, and so on). The literature on microfluidic sensing is very rich. Thousands of individual techniques have been published, with many being just variants of one general idea. The following sections are not intended as a comprehensive review, but rather as a rapid guide to some sensing techniques with important integration potential. We will ignore many techniques that currently employ large external equipment to perform measurements on a microfluidic chip— some of these techniques might however have the possibility of being


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integrated on-chip one day. Wherever an integration path has been identified, it will be mentioned in the text. Just a few relevant examples will be given from published literature, to illustrate various sensing and measurement techniques. To gain a better feeling for the full spectrum of work published in this area, the reader is urged to consult the review literature, which is very rich in itself: an important number of excellent reviews (of which we only name a few) describe recent advances in microfluidic detection technology: [105, 142–153].

2.5.1 Optical Measurements Optics is considered the technology of choice in a large fraction of laboratory measurements. Optical techniques have many advantages—they do not require actual contact with the fluid of interest, but can be performed through the walls of a transparent tube or cell; optical measurement protocols and methods adapted to a wide variety of compounds (of both chemical and biological nature) exist; the sensitivity of optical measurements is generally high, when a sufficient sample volume exists; light can be guided to the region of interest using a variety of techniques, such as optical fibers and waveguides or lens systems. In the following we will not focus on the type of measurement to be performed (i.e., on the chemical compound or the biological entity that needs to be detected), but rather on the optical measurement technique itself. It is important to realize that large constraints exist in microfluidic systems due to the very features that make them interesting in many applications: small dimensions lead to reduced path length for absorbance measurements and to reduced signal in all types of measurements due to the reduced sample amount. There are several classes of optical measurements that are currently being implemented in microfluidic systems. Of these, fluorescence / chemiluminescence detection and absorption spectroscopy are the most interesting from an integrated sensor perspective. Fluorescence is the most widely utilized optical method for molecular sensing in microfluidic systems due to the existence of highly sensitive and selective fluorescent labeling techniques for different biological compounds, or to the existence of predyed particles that can either be used by themselves (e.g., as tracers) or can be functionalized. However, dyes used in fluorescence can be expensive, are sometimes influenced by other chemical properties of the sample, and have a finite lifetime. Labeling of the compounds of interest often requires complicated fluidic manipulations by itself, which makes alternative label-free analysis techniques very attractive in many applications.

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Figure 2.42 The schematic of an epifluorescence microscope, showing the broadband light source, the excitation filter, the dichroic mirror, the microscope objective focusing light on the microfluidic chip, and the detection filter. The fluorescent light is reflected by the dichroic mirror, and then filtered using the detection filter and recorded using ocular optics.

Fluorescent detection requires light of a short wavelength (typically in the UV or blue region of the spectrum), which is absorbed by the molecules of interest and then reemitted at a longer wavelength. In most microfluidic implementations, the light is emitted by a laser source or an LED, which can either be purchased as independent components and integrated in a hybrid system or be incorporated on-chip. A laser light is preferred normally, because of its monochromatic nature and low divergence that makes it easier to focus on the small sample volumes typical of microfluidic systems. LED light sources, on the other hand, represent a lower cost alternative (an issue of concern particularly in disposable applications such as those used in the medical point-of-care field), and can be combined with other optical elements such as waveguides, optical fibers, filters, and integrated lenses to allow better focusing and a narrower light spectrum. Often, the microfluidic chip does not integrate any actual optic devices, but is placed under an epifluorescence microscope and observations are made using standard laboratory optical equipment. The schematic of an epifluorescence microscope used in conjunction with microfluidics is shown in Figure 2.42. A broadband light source is normally used (typically a halogen


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lamp is used, replaced by LED in recent models), which is then filtered using the emission filter to allow only the short wavelengths of interest for fluorescence to reach the sample (most often a highpass or a bandpass filter is used). The filtered light travels through the dichroic mirror and is focused on the sample contained within a microfluidic channel of chamber. Both scattered light and fluorescent light enter the objective, and are filtered by the dichroic mirror and the detection filter. The fluorescence cubes (as the emission-dichroic-detection filter combination is called) are applications and fluorophore-specific. The most widely used fluorophores, which are fluoroscein and rhodamine 6G, fluoresce in the green and the red regions of the spectrum, respectively. While this epifluorescence setup could be interesting in large-scale expensive equipment where the microfluidic chip is only used to reduce sample volumes and increase the speed of the experiment, its large volume makes it prohibitive in most applications where some degree of portability is required. A lot of effort has therefore been guided at integrating at least some of the optical chain on-chip. A pioneering effort has been made in the monolithic integration of optical fluorescent detection in a DNA-analysis device [154]—in that case, an external blue LED light source was used, and the photodetectors were integrated underneath the fluidic channel. In addition to the optical detector, the system included several sample preparation and detection components (liquid injector, mixer, temperature-controlled reaction chamber, electrophoretic separation). The system was realized on a silicon substrate using standard microfabrication techniques; a schematic and a micrograph of the complete system are shown in Figure 2.43. The same group developed the technology further, using parylene channels on a silicon substrate and integrating an interference filter onto the photodetectors to prevent excitation light from affecting the detection [155]. Several similar examples, involving detection of either fluorescence or chemiluminescence using integrated photodetectors, exist in the literature and are reviewed in the literature cited at the beginning of this section. Other efforts where optical detectors have been integrated into a microfluidic system involve fabrication using silicone elastomer (PDMS) using soft lithography molding techniques. A system integrating an optical fiber for guiding the excitation light from an external LED to the detection microfluidic channel, disposable fluidic circuitry, a filter, and a reusable microavalanche photodiode has been demonstrated [156]. Another example of integrated emission-fluorescence-detection system incorporated a blue LED dye positioned on top of a photodiode covered by an interference filter. The resulting optical system was then placed underneath a PDMS microfluidic device to excite and detect fluorescence in a microchannel [157] (Figure

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Figure 2.43 An integrated DNA analysis device with fluorescent detection. Several fluidic components, heaters, an electrophoretic separation cell, and photodetectors were integrated on a single silicon chip. (Reproduced from [154].)

Figure 2.44 A hybrid-integrated microfluidic device with LED emission, inteference filter, and photodetector placed underneath a microfluidic device. Left: The schematic of the device showing the PDMS channel bonded to a glass slide (top), the emitting blue LED dye positioned on a reflective pad, the optical filter and the photodetector. Right: The resulting complete system, with the LED active. (Reproduced from [157], with permission from Elsevier.)

2.44). While the integration of external components in PDMS involves rather complicated manual fabrication steps, it provides a relatively low-cost alternative to monolithic integration approaches, and has the advantage that it can often be realized in a regular laboratory without access to expensive clean room equipment. Absorption spectroscopy is another widely used laboratory technique.


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It relies on the absorption of light by different molecules at specific wavelengths due to coupling of the electromagnetic energy with different rotational or vibrational modes of the molecule. Since the absorption spectra are very specific to the type of molecule, they represent a very useful analytical chemistry technique. Once a compound of interest has been identified and its absorption spectrum measured, the concentration of the compound in a sample can be determined using the Beer-Lambert law by performing the measurement of absorption at one wavelength where the molar absorptivity is known (alternatively, a full spectrum can be recorded, in which case an external spectrometer is typically required). The Beer-Lambert law relates the transmitted light intensity through a sample Φt to the incident light intensity Φi , the molar absorptivity of the compound at the measurement wavelength α, the concentration of the compound in the sample c, and the total path length of the light through the sample l: Φt = Φi 10−αcl

(2.2)

We can see that there is an exponential dependence of the transmitted light intensity on how absorptive the compound is, how concentrated it is, and how long the light travel path is. Usually it is only the last parameter that can be controlled in a microfluidic cell, to bring the total transmitted intensity within the detection range of the photodetector used. Typical lengths that are required for absorption detection in most analytical chemistry applications range from millimeters to several centimeters—standard cells for benchtop instruments are available from 1 mm to 10 cm lengths. Transverse microfluidic channel dimensions, on the other hand, range from a few to hundreds of micrometers—two to three orders of magnitude below the ideal size; microfluidic systems can be used with transverse light absorption only if very concentrated and highly absorptive fluids need to be measured. The length of microfluidic channels, on the other hand, can easily reach several millimeters, which becomes interesting for a wide variety of absorption measurements. It is difficult, however, to channel light along a channel without incurring significant losses due to a lack of collimation, alignment, or absorption in the walls of the channel. The attainable detection limits in microfluidic optical absorbance implementations therefore depend critically on the optimization of the optical chain—effective light guiding, collimation, and collection are required. The majority of published efforts involving absorbance detection in microfluidic systems involve light guiding using one or several optical fibers for coupling light into and out of the microfluidic chip. In some cases, the light is incident transverse to the microfluidic channels via optical fibers brought in

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the proximity of the microfluidic chip [158] or inserted in perpendicular sidechannel fiber guides. External light sources and miniaturized spectrometers are typically connected to the free ends of the light injection and detection fibers. Such optical arrangements result in a very short optical path and relatively poor detection limits; they are applicable in situations where the compound to be measured is in abundance. Alternatively, the light path length can be improved by including some optical components on the actual microfluidic chip. Light from the optical fibers, which are guided using side channels within the chip, can be collimated in-plane using either PDMS lenses [159] or curved air mirrors [160]. Such arrangements, typically realized in PDMS material using soft lithography, can provide significant detection limit enhancements due to the higher signal levels induced by collimation. Stray light can also be eliminated by surrounding the ends of the optical fibers with channels filled with absorptive materials like black ink [159] (Figure 2.45, top). Similar devices have also been realized in silicon and glass technology, using UV-curable epoxy to secure and hermetically seal the fibers (which had the ends in direct contact with the fluid) within channels [140] (Figure 2.45, bottom). Microfabricated UV-transparent silicon oxinitride waveguides have also been integrated on an oxidized silicon substrate and used to guide light from the optical fibers to the microchannels and back [161]. A higher degree of integration can be achieved by integrating a light source, waveguides, and photodiodes into a single microfluidic system. Figure 2.46 shows a recent example [162], integrating an on-chip liquid dye laser, coupled to a set of waveguides and to fluidic channels fabricated in SU8 polymer and incorporating photodiodes on the substrate. Liquid core waveguides can also be integrated with solid-core waveguides to provide improved coupling between the light source and the microfluidic specimen [163].

2.5.2 Chemical Measurements Chemical detection is an enormous field, and covering all possible detection methods is outside the scope of this chapter. Among chemical measurement methods that do not utilize some form of optical detection, we will only provide a brief introduction to electrochemical techniques, and we will focus exclusively on the technology involved in the detection itself, and not on customizations that are specific to the actual compound being detected. Only a few examples will be given; complete literature surveys have been provided in the review articles cited earlier.


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Figure 2.45 Top: Schematic of a PDMS device made of three molded layers and incorporating fiber-guiding channels, flow channel, PDMS lenses for light collimation, and absorptive fiber guards to eliminate stray light. Bottom: Silicon/glass microfluidic spectroscopic flow cell, with input and output fibers sealed using UV-curable epoxy and with the ends in direct contact with the fluid. (Reproduced with permission from [159] and [140] (Copyright 2005, 2010 American Chemical Society), respectively.)

Figure 2.46 The microfluidic chip shown above consists of a light source (liquid dye laser), solid core waveguides, fluidic components, and on-chip photodiodes, and is an example of an optofluidic device that has achieved the highest degree of integration to date. (Reproduced from [162], by permission of The Royal Society of Chemistry.)

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In the previous section, we have studied a number of microfluidic optical detection systems with different degrees of component integration. While some integrated optofluidic devices have been demonstrated, the field is still in its infancy and important technology developments are still required to enable widespread optofluidic on-chip integration. As proof, the majority of the systems reviewed in the previous section employed at least some sort of laboratory-sized external component, such as a microscopy setup or a spectrophotometer. Electrochemical methods provide an alternative detection approach, offering both high sensitivity and compatibility with microfabrication. Capillary electrophoresis chips with integrated injection and electrochemical detection resulting in a complete microanalysis device were manufactured in a pioneering effort more than a decade ago [164]. Since then, hundreds of microfluidic designs incorporating some form of electrochemical (amperometric or conductivity) detection have been reported. The basic design used in microfabricated electrochemical devices relies on more or less complex patterns of electrodes manufactured using different types of deposition processes. The most typical amperometry configuration consists of three electrodes: a reference electrode, a working electrode, and an auxiliary electrode, possibly separated from the microfluidic channel by an ion-selective membrane to improve the specificity of the sensor. An electrical potential difference is applied between the working electrode and the reference electrode (the working electrode is being polarized), and the current flowing between the working electrode and the auxiliary electrode provides the measurement of analyte concentration. By comparison, conductivity detectors (used primarily as ionic detectors, and usually in conjunction with an electrophoretic separation step) are based on a two-electrode design. The electrodes can be in contract with the fluid or separated from it by a dielectric layer; in both cases the electrical measurement is performed in AC mode. A good example of the versatility of microfluidic electrochemical detection is given by a multicomponent sensor for monitoring glucose, lactate, glutamate, and glutamine [165]. Working electrodes were manufactured out of electrodeposited platinum, whereas the reference electrode was manufactured out of Ag/AgCl (electrodeposited Ag, galvanically chlorinated). The counterelectrode was manufactured on the PCB to which the biosensor was attached. The specificity of each electrode was given by a layer of hydrogel in which specific enzymes were immobilized during the crosslinking process. The sensor was completed by attaching the active part to a microfluidic channel that incorporated a mixer on chip. The resulting electrochemical biosensor is shown in Figure 2.47.


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Figure 2.47 A complete biosensor based on electrochemical detection. Left: The complete system, showing the active sensor part, the connection electrodes, and the fluidic channels. Right: The detail of the functionalized biosensor electrodes. (Reproduced from [165], with permission from Elsevier.)

2.5.3 Physical Measurements The physical properties of fluids—namely, viscosity, density, surface tension, and the fluid flow rate within a microfluidic channel, are very important in a number of applications, ranging from the medical field to various industries such as chemical engineering, fossil fuel production and automotive. By comparison with the previous two sections, which included mostly optical and electrochemical techniques, the fluid properties are typically measured using physical principles based on thermal and/or mechanical interactions. Most mechanical measurements in microfluidic devices involve the deflection or the resonance of an elastic component, be it a mechanical beam, a membrane, a tube, or a resonant plate. In order to achieve accurate and reproducible mechanical measurements, a material with stable and well-understood elastic parameters is required. The material should also be impermeable to the fluid of interest (unless, of course, the measurement depends somehow on the swelling of the material). These requirements exclude by default many polymeric materials, which age rapidly with time and have limited chemical compatibility and relatively high affinity for solvents. The materials typically used for building microfluidic devices containing mechanical sensing elements turn out to be, not surprisingly, silicon, glass, and various metals. It is interesting to note that, while numerous reports of novel microfluidic optical and chemical sensing techniques are made every month, the publications attempting to perform measurements of certain physical properties of a fluid in a microfluidic system are much rarer. This is due on one hand

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Figure 2.48 Cross-section through a silicon microtube (left) compared with a human hair (right). The SEM magnification was 750× in both cases. (Reproduced from [167], by permission of The Royal Society of Chemistry.)

to the smaller available market—the vast majority of published microfluidic applications lie in the areas of biology and chemistry, where optical and electrochemical measurement techniques are predominant. On the other hand, this is due to the objective difficulty of manufacturing, integrating, and packaging together MEMS and microfluidics, which requires expertise in multiple domains and access to expensive manufacturing facilities. An original device incorporating a vacuum-packaged suspended microfluidic U-tube was manufactured in silicon together with capacitive sensing electronics [166, 167]. The tube was driven in resonance capacitively and the resonant frequency was correlated with the total tube mass, and hence to the mass of the fluid within the tube and therefore to the density. Figure 2.48 shows a cross-section through the silicon tube used in the device. More interestingly, by flowing fluid through the tube, the out-of-plane oscillation of the tube due to the Coriolis force could be detected, resulting in a measurement of mass flow rate. Recently, it was shown that the damping of the tube is directly related to the viscosity of the fluid inside the tube, resulting in a measurement of both kinematic and dynamic viscosities [168]. More details about the commercial applications of this technology can be found in Chapter 5. A device operating on similar resonant principles, but using a plate oscillating within the fluid of interest instead of a resonant tube, was used to measure both the viscosity and the density of a fluid [169]. The same authors later demonstrated a metallic resonant tube device with a total internal volume of 20 µm, which was used to measure the density of fluids in high-pressure and high-temperature conditions [170]. Microfluidic devices have recently been identified as a very promising


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Figure 2.49 A microfluidic viscometer/rheometer device integrating in a monolithic fashion a differential pressure gauge for stress measurement, a resistive temperature detector, and a microfluidic time-of-flight flow rate sensor. This device is an example of the current state of the art for integration of microfluidics with MEMS mechanical sensors. (Reproduced from [173].)

platform for performing fluid viscometry and rheology measurements [171]. Fluid viscosity sensors based on measurements of the pressure drop along a fluidic capillary have recently been demonstrated in silicon-glass technology, utilizing multiple absolute capacitive pressure gauges distributed along a microfluidic channel [172]. More recently, the integration of a differential pressure gauge with on-chip temperature detectors and thermal time-of-flight flow rate detection has resulted in a fully integrated viscometer. The device was made using SOI technology for manufacturing the piezoresistive pressure gauge membrane as well as the heater and detector bridges. The silicon wafer incorporating the sensor elements was then sandwiched between two micromachined Pyrex wafers, which defined the upper and lower channel walls and contained the fluidic ports [173] (Figure 2.49). Flow rate measurements are important in many scientific and industrial applications involving process control (e.g., analytical chemistry, chemical engineering, combustion engine applications) or flow metering (e.g., biomedical applications such as drug delivery, blood flow, and respiration; gas and fuel meters; air conditioning). Microelectromechanical systems (MEMS) technology has been applied for many years to the microfabrication of flow measurement sensors for either gas or liquid flow [174, 175]. Several efforts have recently been directed at integrating flow measurement sensors within microfluidic systems [4], utilizing techniques as diverse as thermal anemometry, calorimetry and/or time of flight [176–178], differential pressure [179, 180], and the Coriolis effect [181, 182]. The technologies used for manufacturing microfluidic flow sensors range from traditional silicon/glass

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MEMS technology to microfabrication using polymeric materials. Thermal techniques stand out among others through their simplicity in both the fabrication process and subsequent sensor conditioning and measurements [1]. Thermal sensing methods such as anemometry and calorimetry suffer from a number of specific drawbacks: they have some dependence on the fluid’s properties (most notably on viscosity, thermal conductivity, and heat capacity), and they usually have a response that does not depend linearly on flow velocity, therefore requiring calibration and linearization. Time-of-flight techniques have been developed to avoid these issues, with several microfluidic implementations [177, 183]. A stochastic time-of-flight technique has recently been implemented in a microfluidic device, resulting in a flow sensor with a response that does not depend on fluid properties and is linear over close to three orders of magnitude [184].

2.6 Conclusions The overwhelming number of technologies and devices presented in this chapter stand witness to the fast pace at which the field of microfluidics is evolving: the large majority of the referenced articles are no older than a decade. The coming decade will likely be even richer in technology developments as well as in new applications. The capability to integrate multiple functional blocks in a single device will become an essential ingredient for commercial success, as the microfluidics market becomes more competitive and cost-conscious. This chapter reviewed a large number of microfluidic functional building blocks, outlining the underlying operation principles, the manufacturing technique and materials, the performance of the component (whenever data could be found), and the possible drawbacks of the technology. This chapter covered the technologies most often employed for performing flow manipulations in integrated microfluidic systems; both single phase and multiphase approaches were presented. Detection technologies, which tend to be very application-specific, were covered only generically, with a few specific examples. The functional blocks presented in this chapter should enable the reader to have a broad view of the available options and to make educated choices regarding the technologies to employ in a new project.


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3.1 Introduction Microfluidic devices deal with phenomena at scales from one to a few hundred micrometers. At such small scales, the forces that govern the behavior of many systems can be different from what we are used to at the macroscopic scale. It is very useful to develop a physical understanding of the phenomena that are becoming dominant at the microscale; this understanding will allow us to identify and to focus on the most important aspects the fluidic architecture, resulting in a capability to design and optimize fluidic systems efficiently. The main focus of this chapter is the single-phase flow and related phenomenology. While the subject of digital and multiphase microfluidics has become quite fashionable these days and is of obvious importance in a number of applications, a detailed explanation of the phenomena encountered in a multiphase flow is beyond the scope of this modest introduction and would require a separate book. The main aspects are treated nevertheless in enough detail to give the reader the basic background to understand, however superficially, the phenomenology. A certain mathematical background is required to follow the material in this chapter.1 The mathematical apparatus required is, however, kept at a minimum; since this book is focusing on microfluidic system design, we will often avoid intricate mathematical derivations in favor of intuitive explanations, models, and rules of thumb. In several places we will supplement the theoretical material with practical examples, calculations, and estimates. The reader is directed to a number of excellent textbooks and treatises that cover the different aspects discussed in this chapter in much greater 1. The reader should be familiar with multivariable calculus, and should be aware of the notation conventions adopted in the text: Vectors are represented as bold characters. For example, v represents the velocity vector, and v represents its magnitude. Unit vectors along the coordinate axes are represented by a hat, ˆ, y ˆ, ˆ as in x z. Similarly, the direction of any vector r is represented by ˆ r = rr . ∂α =

∂ ∂α

where α can be any variable such as x, y, z, t; ˆ ∂x + y ˆ ∂y + z ˆ∂z , the gradient operator ∇=x ∆ = ∇ · ∇ = ∂x2 + ∂y2 + ∂z2 , the Laplace operator The majority of the derivations and physical examples present in this chapter avoid the use of tensors. While this choice renders the bulk of the material less technical and easier to read by a larger audience, it does impose important limitations to the types of phenomena that can be described in this simplified mathematical frame (this is particularly true of fluid dynamics). Fortunately, in the case of microfluidics, which deals mostly with laminar flows of Newtonian fluids in relatively simple geometries, the limitations prove not to be very stringent.

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detail [1–7]. Some of these texts are very mathematical (which is a prerogative of fluid dynamics); however, they will provide detailed explanations to anyone dealing with microscale phenomena or desiring to go beyond the order-ofmagnitude estimates.

3.2 Diffusion Laws We are all familiar with the phenomenon of diffusion: if we place a drop of red-colored ink in a glass of water (the water being stationary and in thermal equilibrium, such that no convection currents are present), we will initially witness a red blob, which will fade over time until the ink becomes uniformly distributed throughout the water volume, giving it a slightly red shade. The ink molecules (or particles, depending on the type of ink) will diffuse from the high-concentration regions to those of lesser concentration. This phenomenon is due to the Brownian motion: thermally equilibrated fluid molecules move in random directions with velocities given by a temperature-dependent distribution. As they impact a foreign particle, they will transfer some of their momentum to the particle, forcing it to follow a similarly random motion (or to diffuse) through the fluid. The foreign particles will thus move around, eventually reaching every region of the fluid volume. This process is called molecular diffusion. Quite naturally, the number of impacts per unit time and the momentum transfered per impact depend on the size, shape, and mass of the foreign particle (or molecule), as well as on the temperature. This explains why the diffusivity of various chemical species varies greatly with the molecular size (or weight) and also with temperature. A related phenomenon happens in a homogenous fluid when different regions of the fluid are at different temperatures. In this case, the random thermal motion of fluid molecules will cause the “hot” fluid molecules (i.e., molecules that carry more momentum or, equivalently, more energy) to collide with “cold” molecules, transferring some of their momentum and energy over and thus “heating up” the “colder” molecules.2 Eventually, the energy (and temperature) will become uniformly distributed throughout the fluid. This process is called thermal diffusion. While molecular diffusion and heat diffusion seem to be unrelated, they are governed by the same physical equations. Indeed, in both cases we have a flux (of mass in the case of molecular diffusion, and of heat in 2. We use quotes around the terms “hot” and “cold” because their use is slightly improper: normally, hot is associated with higher and cold with lower temperature. However, temperature is a statistical property and therefore has no meaning for a single molecule. What we really mean by a molecule that is “hot” and “cold” is that the molecule originates from a fluid region of higher, or lower, temperature.

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the case of heat diffusion) that originates in the random thermal motion of molecules coupled with a nonequilibrium situation (nonuniform concentration or temperature profiles). In both cases, there is a current (or flux) that develops and is proportional to (and opposes) the local concentration (or temperature) gradient. This is called Fick’s first law for molecular diffusion, and Fourier’s law for heat diffusion. These laws can be mathematically described as: Jm = −Dm ∇c (for molecular diffusion)

(3.1)

Jh = −κ∇T (for heat diffusion)

(3.2)

∂t c = −∇Jm = Dm ∇2 c (for molecular diffusion)

(3.3)

∂t ǫ = −∇Jh = κ∇2 T (for heat diffusion)

(3.4)

where Jm and Jh are the mass and heat currents, respectively, c is the concentration of the foreign species, T is the temperature and Dm , κ are the molecular diffusion coefficient and, respectively, thermal conductivity of the fluid. From here we can obtain mass and energy conservation (or continuity) equations in differential form3 (called Fick’s second law of diffusion):

where ǫ is the fluid’s internal energy density. In deriving these equations, we worked in the linear regime by making the implicit assumption that the diffusion coefficient and thermal conductivity do not depend on temperature or on the spatial coordinates. For small concentration and temperature variations these assumptions are justified, but they may break down in certain extreme situations. A typical example involves a heated fluid, which changes its density (due to thermal dilation) and may also change its phase (by boiling); in such cases the equations above require corrections. We notice that (3.4) involves both the temperature and the fluid’s internal energy. To obtain an equation for the evolution of the temperature field alone, we can relate the two using the fluid’s specific heat: δǫ = ρcp δT . In this case, (3.4) becomes: κ 2 ∂t T = ∇ T = Dh ∇2 T (3.5) ρcp with Dh taking the place of an effective diffusion coefficient. We now can see explicitly the similarity between the mass and heat diffusion equations (3.3) and (3.5). In general, the molecular diffusion coefficient depends on the diffusing chemical species as well as on the solvent. The diffusion coefficient of a 3. For a derivation and physical justification of the continuity equations presented here, we refer the reader to the derivation of the mass conservation equation used in fluid dynamics (Section 3.3.1).

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chemical species in a certain solvent is normally an empirical datum. In the case of particle diffusion, however, the size of the particles is much larger than the molecular length scales and the diffusion coefficient can be calculated exactly. The formula, proven by Einstein in his studies of Brownian motion [8], relates the diffusion coefficient of a colloidal suspension to the temperature T , the radius of the particles r, and the viscosity η of the solvent4 : D=

kB T 6πηr

(3.6)

(here kB is the Boltzmann constant). We notice that, as expected, smaller particles diffuse faster, which can be used in a number of applications, such as the separation of various particles by size in an H-filter. From (3.3) and (3.5) it is apparent that the diffusion coefficient has units length2 of time . Using dimensional analysis, one can obtain a number of orderof-magnitude estimates for diffusion phenomena in particular systems. For example, by dividing the diffusion coefficient by the characteristic length of the system (the length scale over which we have significant concentration or temperature variations), we obtain a quantity with units of velocity—this can be thought of as a species’ “migration” velocity in a concentration gradient v = D/L. Similarly, dividing the square of the characteristic length by the diffusion coefficient gives a quantity with units of time, which is the system’s characteristic diffusion time—an estimate for the time it takes for a species to diffuse across the entire system: τ = L2 /D. Finally, if we multiply the diffusion coefficient by a time, we obtain √ the square of a length. This length is called the “diffusion length,” L = Dτ , and it represents the distance that the species has diffused during time τ . As we will see below, such order-ofmagnitude estimates can be very useful in determining which phenomena are dominant in a certain microfluidic geometry. As a simple example, we can calculate the time required for a fluorescent dye such as Rhodamine 6G to diffuse across a water-filled microchannel (D = 2.8 · 10−6 cm2 /s, L ≈ 100 µm): τ = L2 /D ≈ 35 s. This is a rather long time, hence the need for smart mixing geometries that can enhance molecular diffusion either by reducing channel width or by creating a “wavy” or “chaotic” interface between Rhodamine and water. Micromixers form a topic by themselves in microfluidics.

3.3 Fluid Dynamics Fluid flow in microfluidic systems is governed by the same fundamental equations that describe larger-scale phenomena (mass, momentum, and 4. The detailed definition of viscosity is given Section 3.3.3.

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Figure 3.1 Mass conservation diagram.

energy equations), but the relative importance of the different terms appearing in the equations changes. While in large systems inertial forces are generally responsible for explaining a majority of the phenomenology, in microscale systems it is the viscous forces that become the dominant players. This leads to a number of simplifications in the modeling of liquid flow in microfluidic systems: turbulence, which is an inertial effect, can usually be ignored, leading to a very simple description of flow in microsystems: the Stokes (or creeping) flow. Such flows can both be accurately described mathematically and simulated using finite-element analysis computer software. There are, however, cases where these simplifications break down and other terms in the equations need to be considered. This section will attempt to guide the user through a simplified derivation of the most important fluid dynamics equations and of the criteria that can be used to identify the dominant terms. 3.3.1 Mass Conservation Equation Mass conservation in a fluid requires that the balance of mass entering and exiting a given volume within a given amount of time be equal to the change in the mass contained within that volume. To understand how to frame this intuitive fact in a mathematical form, we will make use of the diagram shown in Figure 3.1. We will consider a volume element, fixed with respect to a stationary frame at coordinates [x, y, z] and having rectangular faces of dimensions [dx, dy, dz], which is in the path of a flowing fluid. The volume of our element will be V = dx dy dz. Let’s assume that the mass current density of fluid, J = ρv, defined as the total mass of fluid transported per unit ˆ direction and depends on the x normal area per unit time, is oriented in the x coordinate as well: J = Jx (x)ˆ x

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In this case, the mass of fluid entering our volume element per unit time dt from the left is Jx (x) dy dz dt, and the mass of fluid exiting it from the right is Jx (x + dx) dy dz dt. The total increase in mass within our volume element is therefore given by: dm = Jx (x) dy dz dt − Jx (x + dx) dy dz dt = −∂x Jx V dt Since the density is defined as the ratio mass m to volume V , we can conclude that dρ + ∂x Jx dt = 0 (3.7) In other words, the spatial variation of the mass current density and the rate of change of density must be balancing each other. In the more general case where J(r, t) and ρ(r, t) are arbitrary vector and scalar fields, respectively, (3.7) acquires the following more general form, called the mass conservation, or continuity equation: ∂t ρ + ∇J = 0 (3.8) We notice that in the case of an incompressible fluid the density field is constant and the above equation reduces simply to: ∇v = 0

(3.9)

It is important to realize that (3.8) applies not only to mass, but also to any other conserved scalar variable, such as electric charge. 3.3.2 Momentum Equation Newton’s second law claims that in an inertial frame, the rate of change in an particle’s momentum is equal to the vector sum of all the forces acting upon the particle: dp X = Fi (3.10) dt i In the case of a flowing fluid, we are dealing with a deformable fluid particle (or volume element) that is advected along a flow line. To apply Newton’s second law, therefore, we must compute the forces and the rate of momentum change for this particular fluid particle, and trace it along its flowline (Figure 3.2). Since we always refer to the same fluid particle, its mass mf p is constant by definition and all change in momentum comes entirely from changes in the velocity field. We will describe the velocity and density of the fluid by a vector field v(r, t) and a scalar field ρ(r, t), respectively, referenced to a fixed inertial frame. The trajectory of the fluid particle we are tracing (i.e., its

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Figure 3.2 Fluid particle being traced along its flow line in a fixed inertial frame.

flow line) can be parametrized in this inertial frame by rf p (t). Flow lines are determined by the velocity field, so we must have: v(rf p (t), t) =

drf p (t) dt

(3.11)

The volume of the fluid particle is Vf p (r, t). It is constrained by the condition that the mass mf p of the particle is constant, leading to: Vf p (r, t) =

mf p ρ(r, t)

(3.12)

The momentum pf p (t) of the fluid particle at any moment in time is then given by: pf p (t) = mf p v(rf p (t), t) (3.13) and the rate of change of momentum can be calculated as: dpf p = mf p (∂t v + (v · ∇)v) dt

(3.14)

where we made use of (3.11) in the derivation. The operator Dt = ∂t + v · ∇

(3.15)

is known by many names, the most common being convective (or advective) derivative, flow derivative, or material derivative. It is applied whenever the rate of change of a particular property of a fluid (in this case its momentum) needs to be computed along the path of the flowing fluid, and is very commonly used in fluid dynamics. The convective derivative accounts for

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Figure 3.3 Forces caused by a pressure gradient in the x direction acting on an infinitesimal fluid volume.

changes in a (vector or scalar) property of the fluid due, on one hand, to time variations in the property field itself, and on the other hand to variations in the property value that occur as a result of the advection of the fluid through the nonuniform property field. Newton’s second law (3.10) as applied to a body of fluid then becomes: X dpf p = mf p Dt v = Fi dt i

(3.16)

By dividing both sides of the equation by the volume Vf p of the fluid particle, and making use of (3.12), we obtain the following expression for Newton’s law, which involves the density and velocity fields (ρ, v), and volume force densities fi : X ρDt v = fi (3.17) i

The most common force densities that we will encounter in the study of single-phase microfluidic flows will come from external pressure gradients, internal friction, gravity, and electromagnetic forces, which will be studied in the next section. In two-phase flows we encounter additional surface forces coming from surface tension effects as well as surface charges, which we will treat in a separate section of this chapter. 3.3.3 Fluid Forces 3.3.3.1 Pressure Gradients

A pressure gradient gives rise to forces pushing the fluid along the gradient. To see how this happens and how pressure forces can be incorporated in Newton’s law (3.17), let’s revert to our simplified view of a small fluid volume element, this time located in a pressure gradient oriented along the x-axis (Figure 3.3).

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Let’s assume the pressure at position x is p(x) and at position x + dx it is p(x + dx). The total force caused by pressure forces on the left side of our volume element is Fx+ = p(x) dy dz, where as the total force on the right side is Fx− = −p(x + dx) dy dz. The resulting force acting on the volume element along the x direction is therefore Fx = (p(x) − p(x + dx)) dy dz = −∂x p dx dy dz, and the corresponding force density can be written as fx =

Fx = −∂x p dx dy dz

In this example, the corresponding forces Fy and Fz are zero, due to the lack of a pressure gradient along those directions. The formula for force density associated with the pressure gradient can then be generalized as: fp = −∇p 3.3.3.2

(3.18)

Internal Friction (Viscosity)

Constitutive parts of a fluid interact with each other by complicated mechanisms that are not always fully understood. Molecular forces, mechanical entanglement between larger molecules, jamming or deformation of droplets in an emulsion, momentum transport by diffusion, and molecular collisions all can account for part of what we commonly call viscosity: the transfer of energy between moving objects or fluid layers and the surrounding fluid, resulting in fluid friction forces. Most commonly, “viscosity” refers to shear viscosity, and quantifies the response of a fluid to an imposed shear rate. Depending on how the fluid reacts, it can be categorized as either Newtonian or non-Newtonian. Newtonian fluids respond linearly, with a shear stress that is proportional with the shear rate over a wide range of imposed rates (the constant coefficient of proportionality between shear stress σ and shear rate magnitude |γ| ˙ being called shear viscosity, η = |σγ| ). Non-Newtonian fluids have complicated, non˙ linear behaviors: they can be, for example, shear-thinning or shear-thickening (in which case the shear viscosity η(|γ|) ˙ decreases or increases with imposed shear rate |γ|), ˙ and may even have constant viscosity plateaus over certain ranges of the shear rate [9]. Examples of Newtonian fluids are most gases and homogenous liquids of low molecular mass, such as water, ethanol, isopropanol, and most common alkanes, as long as system dimensions are large compared to average molecular collision length. Examples of nonNewtonian fluids include complex fluids of macromolecular origin (either polymeric melts or concentrated solutions), foams (like shaving gels and whipped cream), emulsions (like mayonnaise or milk), and suspensions of

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Figure 3.4 Forces caused by shear strain on an infinitesimal fluid volume.

solid or deformable particles (like paints and blood). Many fluids encountered in biological and pharmaceutical applications as well as in the food and cosmetic industries have non-Newtonian characteristics. In addition to viscous response to shear, fluids may also exhibit a viscous response to compression, characterized by the second coefficient of viscosity. Due to the relatively modest compression factors encountered in most microfluidic applications dealing with liquids, second viscosity terms will not be considered in this simplified treatment. However, they need to be introduced in the calculations in acoustics problems dealing with sound absorption of a fluid. To see how the viscous drag term appears in the force balance (3.17), we will refer again to a graphical diagram, shown in Figure 3.4. Let’s consider an infinitesimal fluid volume V = dx dy dz placed in a nonuniform velocity field v(r). We’ll observe the fluid element from a frame moving with the fluid (so that the velocity at the center of the element is zero), and we’ll consider, for simplicity, that the x component of the velocity depends only on the z coordinate, all other velocity components being uniform. This model applies well to flow in long microfluidic channels of constant cross-section, or to flow between parallel plates (Hele-Shaw cell). In this simplified model, the only viscous forces acting on the fluid volume will be on the bottom and top surfaces: Fv (z) = −ˆ xη(z)∂z vx (z) dx dy (3.19) ˆ η(z + dz)∂z vx (z + dz) dx dy Fv (z + dz) = x

(3.20)

and the viscous forces acting on all the other lateral surfaces will be zero due to lack of shear. The resulting force on the fluid volume V will be ˆ (η(z + dz)∂z vx (z + dz) − η(z)∂z vx (z)) dx dy Fv = x

(3.21)

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Figure 3.5 Rheology data and fit of the Carreau model for a non-Newtonian solution of polymer (NIST standard reference material 2490 diluted to 25% in hexadecane).

and the corresponding force density: F η(z + dz)∂z vx (z + dz) − η(z)∂z vx (z) ˆ ˆ ∂z (η∂z vx ) (3.22) =x =x V dz In the case of a Newtonian fluid the viscosity can be considered constant, and (3.22) reduces to: ˆ η(∂z2 vx ) fv = x (3.23) fv =

The slightly more general case, where the x component of the velocity depends on both the z and y coordinates, all other components being uniform, can be immediately derived on similar grounds: ˆ η(∂y2 vx + ∂z2 vx ) fv = x

(3.24)

The non-Newtonian case involves a viscosity coefficient that depends on the characteristics of the flow, and particularly on the magnitude of the shear rate q (3.25) |γ| ˙ = (∂y vx )2 + (∂z vx )2

Equation (3.24) acquires the form:

ˆ (∂y (η(|γ|)∂ fv = x ˙ y vx ) + ∂z (η(|γ|)∂ ˙ z vx ))

(3.26)

The specific dependence η(|γ|) ˙ is called the constitutive equation of the fluid. Several model constitutive equations exist that describe different types of nonNewtonian fluids. A commonly used equation is the Carreau model, which

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can be successfully applied to describe certain polymeric solutions and melts (Figure 3.5): η(|γ|) ˙ = η∞ + (η0 − η∞ )[1 + (λ|γ|) ˙ a]

n−1 a

(3.27)

This equation models a fluid having a Newtonian plateau for small shear rates (η ≈ η0 for |γ| ˙ → 0), and evolving towards a shear-thinning power-law fluid for large shear rates (η(|γ|) ˙ ∼ |γ| ˙ n−1 ). For the case shown in Figure 3.5, the Carreau fit reveals n = 0.7. 3.3.3.3 Body Forces: Gravity and Electromagnetic Forces

Gravity acts throughout the body of the fluid, and its effect on a fluid volume V = dx dy dz is proportional to the mass, and hence the density of the fluid: Fg = ρgdx dy dz

(3.28)

with a corresponding force density f given by: fg = ρg

(3.29)

In the case of a static fluid column not exposed to any other external forces, the gravitational force is normally balanced by the pressure gradient, leading to the familiar equation relating the hydrostatic pressure to depth: p(h) = ρgh

(3.30)

Similarly, if the fluid carries a certain charge distribution ρel (r) and it is exposed to an electric field E(r), there is a Coulomb electrostatic force density on the fluid, given by: fel = ρel E

(3.31)

If there is also a magnetic field present, an additional component appears due to the Lorenz force on the moving charges: fmag = ρel v × E

(3.32)

3.3.4 Navier-Stokes Equation Combining (3.17), (3.18), (3.24), (3.29), (3.31), and (3.32), we obtain the following expression for Newton’s law: ρDt v = fp + fv + fg + fel + fmag

(3.33)

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or, in an expanded form: ∂v + (v · ∇)v) = −∇p + η(∂y2 + ∂z2 )v + ρg + ρel (E + v × B) ∂t (3.34) 2 If the velocity v also depended on x, an additional ∂x term would need to be added to the viscous term on the right-hand side,5 leading to: ρ(

ρ(

∂v + (v · ∇)v) = −∇p + η∆v + ρg + ρel (E + v × B) ∂t

(3.35)

This equation is called the Navier-Stokes equation, and it is one of the most fundamental equations of fluid dynamics. Its nonlinear character, arising from the second term on the left-hand side, makes it very rich in phenomenology: it is this equation that accounts for a complete array of phenomena ranging from flow of thick volcanic lava to the complicated turbulence patterns created in the wake of an airplane. 3.3.4.1

Reynolds Number Re

In many problems the flow pattern in a certain channel geometry or around an obstacle needs to be determined, either experimentally (by performing flow observations), theoretically (by solving the Navier-Stokes equation), or numerically (by modeling it). Most of the time the middle alternative is intimidating, and very rarely a complete solution to the Navier-Stokes equation can be found by theoretical means. It is very instructive, however, to analyze the relative weight of different terms in the equation. As we will see, only a few parameters are actually needed to characterize the flow. This is a particularity arising often in physics: the equations describing a system can be written in terms of variables and differential operators that are dimensionless, all the external parameters describing a system being contained in a few dimensionless ratios. The equations thus become universal, and instead of describing only the specific problem at hand, they actually describe a whole class of phenomena. Fluid dynamics abounds in such dimensionless ratios, which stand witness to the phenomenological richness of this field and quantify the importance of different effects: capillarity, natural convection, relaxation of large polymeric molecules, importance of molecular versus continuous treatment of flow, and so forth. In addition, most of these ratios have names, with historical, physical, and sometimes biblical 5. The derivation of the complete Navier-Stokes equation requires familiarity with tensor calculus, and will therefore be omitted in this simplified treatment. The treatment provided in this section should, however, place the reader in a position to understand the physical origins and appreciate the relevance of the various terms appearing in the Navier-Stokes equation.

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connotations. A sequence of words like “flow at low Reynolds number, but high capillary and Deborah numbers” means something very specific to a chemical engineer or a physicist versed in fluid dynamics, but may intimidate the uninitiated. Nevertheless, such dimensionless ratios are very important since they frame the problem, providing a guide to which terms are important and which terms may be ignored in the equations, and thus highlighting the relevant phenomenology. In the following we will see how to write the NavierStokes equation in a dimensionless form, and from there we will infer that many apparently unrelated flow situations can be characterized by a single parameter, the Reynolds number, which describes the relative importance of inertial versus viscous terms. For identical Reynolds numbers, we obtain identical flow solutions. We will begin by selecting characteristic scales of our system. We will consider first the characteristic length scale Lc for our system, which might be the diameter of a fluidic microchannel, the lateral size of an object suspended in flow, or the size of an orifice. For channels of irregular shape, the length scale is often chosen as the hydraulic diameter dh = 4A/P , where A is the area and P the perimeter of a channel cross-section (we can see that for a circular section the hydraulic and the actual channel diameter coincide, whereas for a square cross-section the hydraulic diameter is equal to the side of the square). We will then pick a characteristic velocity vc , such as the imposed velocity of the fluid at the microchannel entrance, or the flow velocity far from an obstacle. The ratio Lc /vc gives us a characteristic time tc , and the combination ηvc /Lc , which has units of pressure, gives a characteristic pressure pc . If we introduce the dimensionless scaled variables and operators vs = v/vc , ts = t/tc , ps = p/pc , ∇s = Lc ∇, and ∆s = L2c ∆, and for the time being ignore electromagnetic and gravitational terms, the Navier-Stokes equation (3.35) becomes: ρ(

vc ∂vs v2 pc ηvc + c (vs · ∇s )vs ) = − ∇s ps + 2 ∆s vs tc ∂ts Lc Lc Lc

(3.36)

or, equivalently, after simplification: Re (

∂vs + (vs · ∇s )vs ) = −∇s ps + ∆s vs ∂ts

(3.37)

The factor multiplying the left side of the equation Re =

ρvc Lc η

(3.38)

is called the Reynolds number for the problem at hand, and has a remarkable property: in the absence of external forces, it uniquely characterizes the

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solution to the flow equations. In other words, two problems at completely different scales may have identical solutions provided they have equal Reynolds numbers. Flow of water (η=1 mPa s) in a microchannel (Lc =10 µm) is therefore formally equivalent, under similar velocity conditions and upon appropriate scaling of the variables, to drinking syrup (η=1,000 mPa s) using a thick straw (Lc =1 cm). It is interesting to describe the flow characteristics for different Reynolds number regimes, as the types of flow encountered can be very diverse. For Re < 1, the velocity field is time-independent, fully laminar, and completely reversible—reversible recirculation regions may exist at some corners6 but not in the wake of flow obstacles. As the Reynolds number grows to Re ≈ 10, we start to observe some recirculation in the form of eddies or vortex pairs behind obstacles; hence, the flow becomes irreversible but still maintains its laminar characteristics. At Re ≈ 50 the vortices start to break off in an alternating fashion and are advected by the fluid downstream of an obstacle. In this case the flow becomes unsteady, with a somewhat cyclic time dependence. As Re increases further, there are regions of turbulence that develop behind the obstacle and the flow starts to become chaotic. As an exercise we’ll estimate the Reynolds number for a situation common to microfluidic applications. Consider a microchannel of rectangular cross-section 20×100 µm2 , hence a characteristic length (considered to be the hydraulic diameter) of dh ≈ 30 µm. Suppose the fluid flowing through the microchannel is water (ρ = 103 kg/m3 , η = 10−3 Pa·s), at a rate of 1 µl/min. Assuming a flat velocity profile (an approximate description providing a good order of magnitude estimation), this corresponds to a linear velocity of ≈ 8.3 × 10−3 m/s leading to a Reynolds number of Re ≈ 0.25. Such low Reynolds numbers are typical of many microfluidic applications, characterized by very simple and reproducible flow profiles. 3.3.5 Stokes (Creeping) Flow As we have seen above, in usual microfluidic situations one deals with relatively small Reynolds numbers, Re < 1. In these conditions, we can ignore the left-hand side of (3.37) and we obtain a very simple equation describing the flow: ∇s ps = ∆s vs (3.39) which is called the Stokes equation. If we revert back to the physical variables p, v, we obtain the following form for the Stokes equation: p c Lc ∇p = ∆v = η∆v (3.40) vc 6. More details about this type of recirculation pattern are given in Figure 3.8.

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It is interesting to write down the specific form of the Laplace operator ∆ in different coordinate systems. This will prove useful when analyzing flow problems in geometries with specific symmetries. In Cartesian coordinates (x,y,z), it takes the simple form: ∆ = ∂x2 + ∂y2 + ∂z2

(3.41)

whereas in cylindrical coordinates (r,θ,z), it becomes: ∆=

1 1 ∂r (r∂r ) + 2 ∂θ2 + ∂z2 r r

(3.42)

and in spherical coordinates (r,θ,φ): ∆=

1 1 1 ∂r (r 2 ∂r ) + 2 ∂θ (sin θ∂θ ) + 2 2 ∂φ2 r2 r sin θ r sin θ

(3.43)

To find a solution to this equation in a specific geometry, one also needs to apply specific boundary conditions. If a stationary solid wall is present, the fluid velocity at the wall has to vanish (v|wall = 0). The general applicability of this no-slip boundary condition has been debated by some scientists, and may not hold true when flows are examined at molecular scales or when gas bubbles are trapped at the wall due to surface rugosity. Nevertheless, it applies in practically every case of practical interest, and one has to imagine quite exotic experiments to even approach the limits of its validity. In the case where a fluid surface is “free”, that is in contact with a gas or with another medium of negligible viscosity (and which, consequently, cannot impose a shear stress on the fluid), the normal derivative of the velocity at the free surface has to vanish: (ˆ n · ∇)v|surf ace = 0, where n ˆ denotes the unit vector normal to the surface (this can be seen, for example, by examining (3.19) when the left-hand side vanishes). The Stokes equation, in the form (3.39) or (3.40), can be solved exactly for a number of simple geometries, including flow between parallel plates (Hele-Shaw cell), in a straight cylindrical capillary, and the flow around a spherical obstacle. The flow in a rectangular geometry cannot be obtained in closed form, but can be calculated as the summation of an infinite series. In the following we will derive solutions to the simplest flow problems, and will only mention the relevant formulas for problems having important applications but requiring more elaborate calculations. 3.3.5.1 Flow in a Parallel Plate Geometry (Hele-Shaw Cell)

Imagine the situation illustrated in Figure 3.6, where unidirectional flow of an incompressible fluid is imposed between two infinite parallel plates. In

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Figure 3.6 The parallel plate (Hele-Shaw) geometry.

steady state, the velocity vector will be oriented in the x ˆ direction and will not depend on the x or y coordinates, which is a consequence of the symmetries present in the system and of the requirements of mass conservation. The pressure, on the other hand, will only depend on the x coordinate (any y dependence is excluded by the translational symmetry, and a z dependence could only be balanced by a viscous term; given that the z velocity component is null, there cannot be a viscous component in the z direction). Assuming ˆ ∂x p (i.e., pressure varies linearly with a constant pressure gradient ∇p = x the x coordinate and, in our convention, decreases—therefore, the pressure derivative ∂x p is typically negative), the Stokes equation (3.40) in Cartesian coordinates becomes: ∂x p ∂z2 vx = (3.44) η with the no-slip boundary conditions vx |z=−g/2 = vx |z=g/2 = 0 (here g is assumed to be the gap between the plates, and the z origin is taken midway between the plates). The solution that satisfies all the above constraints is a parabola ∂x p 2 g2 vx (z) = (z − ) (3.45) 2η 4 We see that the flow profile is parabolic, having the maximum velocity at the middle of the gap and zero velocity at the walls. It is always interesting to calculate the average velocity vx,avg in a certain geometry given a driving force such as a pressure gradient. This is particularly important

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for estimating time of flight and for making certain hydraulic back-of-theenvelope calculations (a typical example is the calculation of the maximum pressure required at the inlet to push fluid at a given rate through a given microfluidic geometry). In the case of the Hele-Shaw geometry, the average velocity is given by: vx,avg

1 = g

Z

g/2

−g/2

vx (z)dz = −

∇p 2 g 12η

(3.46)

We notice the quadratic dependence of the average velocity on the gap size g, which denotes a high sensitivity to the device geometry (halving the gap will reduce the average velocity by a factor of 4). This is rather typical of microfluidic systems, as we will see: the smallest geometrical dimension has a very pronounced effect on the hydraulic behavior of a system. This effect is even more pronounced in situations where all the dimensions are constrained, such as inside a circular or rectangular channel, and is widely used in hydraulic systems of all sizes, where restrictors (or chokes) of specific geometries are often used to control, or limit, the fluid flow. 3.3.5.2 Flow in a Circular Channel

An infinite circular channel of radius R has cylindrical symmetry; hence, it helps to use the Stokes equation (3.40) with the Laplace operator expressed in cylindrical coordinates (3.42). If we assume the channel to be sufficiently long, end effects can be ignored and one can assume translational symmetry in the z direction: in this case all dependence of the velocity on the z coordinate disappears. The same applies for the θ dependence due to the rotational symmetry, in which case the Stokes equation takes the simple form: ∂z p 1 ∂r (r∂r v) = r η

(3.47)

and its general solution is of the form: v(r) = A + B ln r + Cr 2

(3.48)

∂z p 4η , B = 0 (due to the requirement that v(0) remains finite) and ∂z p 2 −R 4η due to the no-slip boundary condition at r = R: v(R) = 0. We

where C =

A= can therefore write:

v(r) =

∂z p 2 (r − R2 ) 4η

(3.49)

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This expression is somewhat similar to that obtained for the Hele-Shaw cell, in the sense that flow also has a parabolic profile (the flow profile in this case is called Poiseuille flow). We can also calculate the average velocity vavg : vavg

1 = πR2

Z

0

R

2πrv(r)dr = −

∂z pR2 8η

(3.50)

and we notice again the dependence of the average velocity on the square of the radius. We also notice that v(0) = 2vavg (i.e., the maximum flow velocity, which occurs at the center of the channel, is precisely twice higher than the average velocity, which would be the fluid velocity in case all fluid circulated at the same speed). If we look at the volumetric flow rate Φ, through a circular channel, we obtain the following expression: Φ = πR2 vavg = −

πR4 π∂z pR4 = (pin − pout ) 8η 8ηl

(3.51)

By equivalence with electrical resistance circuits, where the value of the electrical resistance R is defined as the ratio between the voltage difference between the two ends of a resistor (∆U = Uin − Uout ) and the current intesity I through the resistor: R = ∆U I , we can define a hydraulic (or hydrodynamic) resistance of a channel ρ as the ratio between the pressure difference between the inlet and the outlet of the channel (∆p = pin − pout ) and the flow rate Φ: ρ=

∆p Φ

(3.52)

By using (3.51) we therefore obtain for the hydrodynamic resistance of a circular channel: 8ηl (3.53) ρ= πR4 In this case we notice an inverse fourth power dependence on the channel radius or diameter, which has enormous implications in microfluidic design: a system designed to operate at a given flow rate Φ needs to be capable to withhold an entry pressure of order ρΦ. Halving the radius of the fluid channel increases this pressure sixteen-fold. Since in most microfluidic systems there is an obvious interest to reduce the overall system volume (e.g., to use less reagent, smaller assays), there is a tendency to reduce the channel diameters. Due to the inverse fourth power dependence of the hydrodynamic resistance on diameter, this may lead to a dramatic increase in hydrodynamic resistance, which will in turn lead to inlet pressures beyond the capability of the system and hence to catastrophic failure. Most often this translates in device failure by fracturing and delamination in leaks at the syringe piston and at tubing

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connections or in syringe or pump failure. Such outcomes are best avoided by a correct choice of channel dimensions for the application at hand. We will come back to the issues of hydrodynamic resistance and electrical modeling of fluidic circuits in Section 3.3.6. 3.3.5.3 Flow in a Rectangular Channel

The most commonly encountered channel geometry in microfluidics is rectangular. Ironically, the solution to the Stokes equation for a rectangular channel of width w, height h < w, and length l is quite complicated to derive and does not have an analytical form (it can only be calculated as the summation of a Fourier series); the mathematical exercise required for this derivation is beyond the scope of this text. One therefore needs to employ various approximations to derive the hydrodynamic resistance of a rectangular channel. For aspect ratios close to 1 ( wh ≈ 1) we can estimate the hydrodynamic resistance using the circular chanel expression for Poiseuille flow (3.53) and assuming the diameter of the equivalent circular channel to be roughly equal equal to the hydraulic diameter: D ≈ Dh = 4A/P = 2wh/(w + h). This estimate is reasonably accurate (within 20%) for the square section, but the error increases drastically as the aspect ratio deviates from 1. A better approximation is given by: ρ≈

12ηl − 0.63 wh )

h3 w(1

(3.54)

which has an error of order 10% for w = h, and is accurate to better than 1% for h < 0.7w [10]. 3.3.5.4 Entry and Exit Effects

The low Reynolds number flow described earlier was always considered in infinite length channels, sufficiently far from the entry and exit orifices. This assumption allowed us to consider translational symmetry in one direction and to derive a number of conclusions regarding the particular structure of the flow. Close to the entry and exit zones, however, this assumption does not hold—here the flow does not have translational symmetry, and the velocity vectors are not necessarily parallel to the channel walls. These regions usually extend several hydraulic diameters away from the entry and exit ports of the channel. Here the pressure gradient does not obey the same equations as in the middle of the channel; this has significant implications for very short channels

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whose hydrodynamic resistance cannot be calculated using (3.53) and (3.54). In most practical situations, the length of the channel is much larger than the other dimensions, and the corrections due to the entry and exit regions are insignificant. 3.3.5.5

Hydrodynamic (Taylor) Dispersion

We have seen above that in a circular channel the Poiseuille flow profile is parabolic (3.49) with the maximum velocity vmax = 2U at the center of the channel and zero velocity at the walls (U is the imposed average velocity). If we place a plug of dyed fluid at the entry of such a channel, what will be the time evolution of the concentration profile? We will initially assume that there is no diffusion, D = 0. In this case, the dye plug will assume a parabolic shape, as the dye is advected by the flow (Figure 3.7(a)). In this case, the fluid never mixes fully, but rather the plug gets streched more with increasing time. A concentration sensor placed on the wall of such a channel will never detect a concentration change, since the fluid velocity at the wall is 0. If we now consider a finite diffusion coefficient, the evolution of the plug profile becomes that shown in Figure 3.7(b). In this case, the plug stretches as before, but in addition to that, its concentration profile gets blurred by diffusion. The diffusion does not only occur along the channel, but also across it: at the head of the plug, there will be diffusion from the center of the channel to the walls (maximum concentration occurs at the tip of the parabolic profile), whereas at the end of the plug, there will be diffusion from the walls toward the center of the channel. These transverse diffusive processes tend to smooth out the radial concentration gradients, resulting in a plug with an axial concentration profile given by an enhanced effective coefficient of diffusion (Deff ), which results from the combination of Poiseuille stretching and radial diffusion. The enhanced diffusion coefficient was first calculated by Taylor, hence the name Taylor dispersion, and the result was later improved by others. Taylor’s expression is given by: Deff ≈

r2U 2 48D

(3.55)

This formula has some interesting aspects: first, we notice that the effective diffusion coefficient is smaller in narrower channels; we also notice that there is an inverse dependence on the actual diffusion coefficient D: if diffusion is very strong, convection will not play a large role in the dispersion process. It is important to understand the limits within which the above formula is valid. It clearly only applies to situations where the diffusion time across the channel is

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shorter than the advection time along the channel (r 2 /D ≪ L/U )—otherwise we would find ourselves in the regime represented in Figure 3.7(a) where diffusion does not really play a role (in the above expression, L is the distance along the channel where we are interested in the concentration profile; in practice, this could be the position of a concentration sensor). There is also a lower limit to the diffusion time across the channel for this expression to be valid: the diffusion time across the channel must be slower than advection time corresponding to the channel radius (r 2 /D ≫ a/U ) or else convection would play no role in the axial spreading of the concentration profile. Within these limits, however, (3.55) provides a very useful approximation to the concentration profile as a function of time, and can be used to estimate the effective resolution of microfluidic sensors and of instruments dealing with sample flowing through long capillaries such as liquid chromatographs. It is also interesting to look at the case of a noncircular channel. A high aspect ratio rectangular geometry, quite common in microfluidic systems, poses an interesting problem: given that there are two characteristic lengths for the channel (the width w and the height h < w), which one should be used instead of r in (3.55)? It turns out that the relevant length scale in this case is given by w, the larger of the two dimensions; indeed, the diffusion along the large channel dimension is the slowest, and therefore it controls the convection-diffusion mixing processes in the long time limit [11]. The equations derived in this section can be applied identically to the spreading of heat pulses, for example, based on the equivalence between (3.3) and (3.5). In the case of heat, however, additional issues arise from the significant heat conductivity of the walls: most of the time, they conduct heat better than the fluid itself, so there is a strong energy “leak” through the channel walls. While approximate solutions can be found for the evolution of a heat pulse in proximity of highly thermally conductive walls, often the best way to arrive at a good order-of-magnitude estimate is by performing numerical simulations. 3.3.5.6 Dead Volumes and Recirculating Flows

One of the important parameters of microfluidic systems is the cleanup time. It is surprising that, despite their small dead volume, microfluidics systems sometimes can have very long reaction times to a change in the environment. This slow response may be caused by a number of things, but very often it is due to the laminar nature of the flow combined with the slow diffusion of various chemical species (or of heat). As we have seen in Section 3.3.5.5, the fluid velocity at the center of a circular microchannel is the highest, whereas the fluid at the walls is

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a

b Figure 3.7 Hydrodynamic dispersion in a circular channel in the absence of diffusion (a) and in presence of diffusion (b). The initial injection profile is that of a plug (left side); the subsequent shapes represent a sketch of the concentration profile evolution with time.

stationary. In case a sensing element is placed on the wall of such a channel and a certain change in the sensed fluid parameter occurs abruptly at the entry point of the microfluidic system, the change will be initially registered at the center of the channel, where it is rapidly advected by the flow. The sensing element will start feeling a change only after a certain time, which depends on the volumetric flow rate Φ, the parameter’s diffusion coefficient, and the microfluidic geometry upstream of the sensor (which govern the Taylor dispersion profile). Ideally, to have a sharp response to a change in the environment, a sensor is best placed in the middle of the channel (e.g., suspended on a membrane). This is particularly relevant for thermal sensors such as anemometers. In addition to the spreading of concentration and temperature profiles due to diffusion and Taylor dispersion, we have other mechanisms that can significantly affect the time response of a microfluidic system. In particular, regions of stationary or recirculating flow will severely limit the time response. Since these regions are only coupled diffusively with the “flowing” regions, they react very slowly to changes in parameter concentrations: indeed, a chemical compound will take a long time to diffuse into, and out of, such stationary regions (we have seen an estimate of diffusion time for a typical chemical compound across a 100 µm channel in Section 3.2). It is therefore very important to design the microfluidic system with the smallest dead volumes possible, particularly in regions where concentration variations are expected. External fittings should therefore be avoided as much as possible, with the implications that sample preparation and measurement

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Figure 3.8 In two dimensions, corner eddies exist at all corners where 2α < 146o . Since the fluid in the corner region only interacts by diffusion with the main stream, it leads to a very slow response to changes in concentration. Wide-angle corners as well as rounded corners, by contrast, do not show this behavior.

functions should be integrated monolithically wherever possible. Square channel corners should also be avoided. Corners with the opening angles smaller than approximately 146o (Figure 3.8) always develop eddies (at least in two dimensions) [12], which are diffusively coupled with the main stream. In most situations these should be avoided, as they can significantly reduce the response time for sensing applications. One of the reasons why microfluidic systems are sometimes designed with square corners is convenience. Most modern CAD software packages offer ready-to-use round-corner and tapered connection primitives, and these should be used wherever possible to avoid recirculating eddies. It is, however, useful to note that corner eddies have been put to good use in particle trapping systems; in these cases, cells or colloids can be trapped within the eddies, allowing a number of analyses to be performed that might have been difficult to achieve if the cell or particle were rapidly flowing. It is also important to point out that eddies may be inhibited in flat Hele-Shaw-type geometries due to the additional friction coming from confinement. Surprisingly, dead volumes play a much smaller role in the case of heat diffusion: indeed, dead volumes being typically bounded by solid surfaces with high thermal conductivity, the heat diffusing into such volumes will be dissipated rapidly into the external walls.

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3.3.5.7


Viscous Drag on a Spherical Particle

A flow situation that appears quite often in microfluidics, but where we will avoid the mathematical derivation and only give the result, is that of a solid particle moving in a fluid under the influence of an external force. In this case the fluid creates a drag force on the particle, which depends on the velocity and the size of the particle—the faster the particle moves through the fluid and the higher its diameter, the higher the drag force. The particle will accelerate until the drag force equals the external driving force, the terminal velocity that the particle reaches being determined by the equilibrium condition (zero net force on the particle). The drag force can be calculated from Stokes’ equation with no-flow boundary conditions on the surface of the particle (solid wall), and its expression is: Fdrag = 6πηrv (3.56) where r is the radius of the solid particle, and v its velocity relative to the stationary reference frame of the infinite fluid; the drag force naturally opposes the movement of the particle. The movement of the particle will also create circulation in the fluid on a local scale. The above results apply only to the case where the size of the particle is significantly smaller than the distance to the nearest wall, but provide useful order-of-magnitude estimates of drag forces in most practical situations. To give an example, we’ll calculate the settling velocity of a colloidal suspension of spherical particles of diameter 5 µm in oil (ρ1 = 700 kg/m3 , η1 =10 mPa s). We will assume that the particles are made of polystyrene and are density matched to water (ρ2 = 1,000 kg/m3 ), which is the case for most commercial colloids. The net external force on the particles will be the difference between their gravity and the buoyancy force: F = (ρ2 − ρ1 )g 43 πr 3 . Equating this to the drag force (3.56) allows us to derive the terminal velocity: 2(ρ2 − ρ1 )gr 2 (3.57) vt = 9η For the practical case at hand, we obtain: vt ≈1.6 µm/s. This may seem slow, but in a microchannel that is only 10 µm deep, it will only take a few of seconds for a particle to settle. This is why colloidal particles are better suited as tracers when they are small and as well-density-matched to the fluid as possible. As we have seen in (3.6), the diffusion coefficient for a colloidal particle can be exactly inferred from the particle dimensions and the viscosity of the fluid. We notice the same combination 6πηr appearing in both (3.6) and (3.56). This should not surprise, as both equations rely on similar underlying physics: the same molecular collisions that are responsible for

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random Brownian motion, and hence for diffusion, are also responsible of the drag force in a fluid. This intuitive realization forms the basis of the fluctuation-dissipation theorem, which provides the mathematical framework for deriving (3.6). Given this realization, one may ask the question whether it is possible for the molecular collisions to be so strong as to maintain a colloidal particle in suspension despite the force of gravity. To answer it, we will make some simple estimates: consider a system of height h. The time required for the particle to diffuse the height h is given by τd = h2 /D = 6πηrh2 /kB T . The time required for the particle to settle under the influence of gravity alone is τs = h/vt = 9ηh/(2(ρ2 − ρ1 )gr 2 ). The particle will remain in suspension if τd ≪ τs , which an easy derivation shows to be equivalent to kB T ≫ (ρ2 − ρ1 )r 3 gh. In other words, there will not be any noticeable particle settling if the thermal energy is larger than the potential energy gain obtained by settling. This is a rather intuitive result, but it is very useful in many applications. In particular, it gives an upper limit to the radius of a particle that can be used as a tracer, rmax ≈ kB T /[(ρ2 − ρ1 )gh]. In the case considered above, the maximal particle radius is rmax ≈ 250 nm—this is difficult to use as a tracer unless the particle is dyed using a fluorescent dye, its diameter being close to the resolution limit of the typical optical microscope. If the particle moving through the fluid is not solid, but liquid (i.e., a droplet of an immiscible liquid of density ρ′ and viscosity η ′ ), the expression for the drag force changes. Indeed, in this case the no-flow boundary condition does not hold anymore, and there is circulation that develops inside the droplet itself. The drag force becomes in this case:

Fdrag = 2πηrv

2η + 3η ′ η + η′

(3.58)

This expression is called the Hadamard-Rybczynski formula, and it reduces to (3.56) when η ′ → ∞, which corresponds to the complete rigidification of the droplet. In the above expression it is assumed that the liquid droplet always remains spherical (which is equivalent to assuming an infinite surface tension between the two liquids), and that the interface between the two liquids is clean (i.e., there are no particles or surfactant molecules present at the meniscus, which might rigidify it). Since in the large majority of cases different surfactants are used to stabilize emulsions and assure droplet stability, the above expression only holds approximately in most practical cases. We notice that we can also set η ′ = 0, in which case we obtain the drag force corresponding to a gas bubble.

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3.3.6 Equivalence Between Fluidic and Electrical Circuits Fluidic systems can often be analyzed using equations emerging from the electrical domain. This is particularly true when dealing with incompressible fluids, in which case the charge conservation equations from electricity (e.g., Kirchhoff’s first law) become equivalent to mass (or volume) conservation equations in fluidics. Ohm’s law, as we have seen in (3.52), also has an immediate equivalent in fluidics, where we define the term of hydrodynamic resistance to relate the pressure drop along a channel to the flow rate. Any elastic deformations that lead to an increase in the volume of fluid ∆V , such as due to the bending of channel walls under the action of a pressure difference ∆p, can be seen as the equivalent of electrical capacitances, where charges accumulate when a voltage difference is applied. We will use χ = ∆V ∆p as the symbol for such a hydrodynamic capacitance. In fact, we can make the following identifications:

Equivalence Table Microfluidic Devices p – pressure V – volume Φ – flow rate ρ – hydrodynamic resistance χ – hydrodynamic capacitance

Electrical Circuits U – potential Q – charge I – current R – electrical resistance C – electrical capacitance

This analogy can be extremely useful, particularly when complicated microfluidic networks need to be designed: many electrical circuit simulation packages exist that can be used to calculate flow rates in microfluidic channels and their time transients. The best-known simulator is SPICE, which originated in the early 1970s as an open-source project at the University of California at Berkeley. It still exists as a stand-alone open-source application [13], and it also forms the basis of many commercial simulator products. In addition to allowing numerical modeling, this analogy allows many established theoretical results from electrical circuit theory to be immediately transferred into the fluidic domain. In the following sections we will see how to use such analogies in practice to derive useful information about the behavior of fluidic systems. 3.3.6.1

Viscous Pressure Drop and the Hydrodynamic Resistance

As we have seen in Section 3.3.5, laminar flow in the low Reynolds number limit leads to fully reversible flow patterns and to a pressure drop ∆p that depends linearly on the imposed volumetric flow rate Φ: ∆p = ρΦ

(3.59)

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The proportionality constant was denoted by ρ, and it effectively plays the role of a hydrodynamic resistance. ρ depends on the channel geometry and dimensions (3.52), and (3.54), and it is generally dominated by the smallest channel dimension. An interesting aspect that is often ignored in fluidic simulations is that, under the influence of pressure, channels may change their geometry (see Section 3.3.6.2) resulting in a change of the corresponding hydrodynamic resistance ρ(p), which will normally decrease as the channel inflates. Such effects may not be trivial and do not have an immediate equivalent in electricity; they may introduce nonlinearities that can prove to be critical in understanding a system’s behavior. 3.3.6.2 Elastic Deformation and the Hydrodynamic Capacitance

Microfluidic channels can be fabricated in a variety of materials having very diverse mechanical properties: metal, plastics, elastomers, glass, and silicon (to name a few). The elastic moduli of such materials can vary by over five orders of magnitude between PDMS (Young’s modulus between 400 and 900 kPa depending on the mixing ratio and curing parameters) and silicon (Young’s modulus 185 GPa). The ductility of the different materials used in microfluidics also presents great variation—glass is highly brittle, whereas plastics and metals (such as nickel) can be extremely ductile. Many microfluidic devices are laminated structures— often pressure failure occurs by delamination, particularly in plastic and PDMS devices. The bonding strengths between different microfluidic components depends on the technique used (adhesive bonding between certain materials can be quite poor, whereas anodic bonding in silicon-glass devices and diffusion bonding in metallic devices can exceed the intrinsic material strength). These properties need to be taken into account when designing a microfluidic device that is meant to work under high internal pressure or with significant pressure differentials. Pressure compliance is the property of an object to yield to pressure without disruption. Normally, a microfluidic channels or device will deform under the action of pressure, to the point that elastic stresses in the material will balance the stresses created by the internal pressure. This deformation can be elastic (in which case the channel comes back to its original shape once the pressure is removed) or plastic (in which case the channel remains somewhat deformed). When pressure stresses are increased beyond a limit, the device will fail and burst. Several mechanisms can be responsible for pressure failure: for brittle materials, pressure typically occurs through crack propagation from a point of high stress concentration or from a material flaw.

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The fracturing of the device in this case happens quasi-instantaneously. For ductile materials, the device yields by progressive deformation, which is a relatively slow process. The best way to predict the maximum pressure that a device will be able to hold is by performing a numerical analysis of the stresses and comparing with the published values of the material’s yield stress. In the elastic limit, however, pressure compliance leads to some interesting phenomena that can be modeled using simple electrical analogs. Imagine a microfluidic cavity that deforms under the action of internal pressure. The increase V (p) in the volume of the cavity (Figure 3.9) will depend on the pressure p according to an equation of the type: V (p) = χp

(3.60)

(we have considered the elastic deformation regime, where dependence of volume on pressure is linear—χ is therefore a fixed proportionality constant whose value can be calculated knowing the geometry of the device and the materials’ properties). As we have seen in (3.52), microfluidic channels can be modeled by analogy with electrical resistances, where the role of the voltage is taken by pressure p, and that of the current by the volumetric flow rate Φ. By the same analogy, the role of the electrical charge is taken by the fluid volume V , in which case (3.60) becomes the definition of a capacitance: χ=

V (p) p

(3.61)

Later on we will see how to use such electrical analogies to predict the behavior of microfluidic systems, particularly their response to periodic pressure pulses (as produced by certain micropumps), and to pressure transients in general. 3.3.6.3

Kirchhoff’s Laws

For microfluidic systems operating with fluids in the incompressible regime, which is most often the case, we can derive equivalents of Kirchhoff’s laws from electricity. The mass conservation law needs to be satisfied at any microfluidic junction—the sum of the flow rates going in and out of a node needs therefore to be 0 (we are using the sign convention that flow going into a node carries a + sign, and flow going out a − sign). This is the fluidic equivalent of the first Kirchhoff law. The second law states simply that the sum of all pressure drops in a closed fluidic circuit loop needs to be 0. This is a natural consequence of unicity of the pressure value.

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Figure 3.9 The elastic deformation of a microfluidic cavity as a function of pressure. The expansion of the microfluidic cavity leads to an increase in its volume V (p) that is proportional (in the small deformation limit) to the pressure p.

Figure 3.10 Representation of a portion of a microfluidic circuit: flow rates in the various channels are denoted by Φi , i = 1 − 8, the nodes are denoted by the letters A–D and the pressure drops corresponding to each channel by ∆pij where i and j represent the end nodes of the channel.

Figure 3.10 represents an example of fluidic circuit having four external ports, four nodes, and one loop. Kirchhoff’s laws can be written for the nodes A, B, C, and D (first law) and for the loop ABCD (second law): Φ1 − Φ7 − Φ5 = 0 Φ2 + Φ7 − Φ8 = 0 Φ8 + Φ3 − Φ6 = 0 Φ6 + Φ5 − Φ4 = 0 ∆pAB + ∆pBC + ∆pCD + ∆pDA = 0

for node A for node B for node C for node D for loop ABCD

(3.62) (3.63) (3.64) (3.65) (3.66)

A consequence of Kirchhoff’s laws are the rules for calculating the

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equivalent resistance (or capacitance) of a network of resistors (or capacitors). Since the laws that apply in fluidic circuits are exactly the same as those in electrical circuits, we can use exactly the same rules to calculate the equivalent resistance (capacitance). For hydrodynamic resistors ρ1 , ρ2 in series, the equivalent resistance is given by ρ = ρ1 + ρ2 . If ρ1 and ρ2 are connected in parallel, the equivalent resistance is given by ρ = ρ1 ρ2 /(ρ1 + ρ2 ). For hydrodynamic capacitors χ1 and χ2 in series, we have the equivalent χ = χ1 χ2 /(χ1 + χ2 ); for the same capacitors in parallel, the equivalent capacitance is given by χ = χ1 + χ2 . 3.3.6.4

Periodic Behavior and Time Transients

In addition to simple steady-state situations, electrical models for microfluidic circuits can be used to analyze more complex, periodic, and transient behavior. We will take as an example the case of a fluidic pump that has pulsating pressure output (this is a common occurrence for electrostatically driven, peristaltic, and piezoelectric pumps). The pulsations in flow may be undesirable for certain microfluidic sensors or experiments. In this case, we would like to design a microfluidic dampening mechanism, that can be connected to the pump and will deliver uniform pressure to the microfluidic system. This is the equivalent of introducing a lowpass filter to attenuate the ripple of a noisy voltage source that needs to supply an electrical system. The corresponding fluidic and electrical circuits are shown in Figure 3.11. In the fluidic circuit, the capacitor is implemented as a diaphragm that deflects under differential pressure. To analyze the two circuits, we will consider a periodic signal at the source, V (t) = V eiωt (p(t) = peiωt ), and we will look at the voltage (pressure) across the system, represented as a resistor Rs (ρs ). The system is essentially a voltage (pressure) divider, and the voltage (pressure) across the system is given by: R ) R + Rs /(1 + iωCRs ) ρ ps (t) = peiωt (1 − ) ρ + ρs /(1 + iωχρs )

Vs (t) = V eiωt (1 −

(3.67) (3.68)

We notice that for ω → ∞, Vs → 0 (ps → 0), which means that indeed the circuit cuts off the high frequencies. For ω = 0 we obtain the regular expression for a voltage divider: Vs = V Rs /(R + Rs ) (and, respectively, ps = pρs /(ρ + ρs )). An interesting question to ask is what happens if suddenly the pump is removed from the circuit or stopped. This may happen if we want,

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b

Figure 3.11 The equivalence between an electric and a fluidic circuit is seen in the implementation of an RC lowpass filter: in the electrical case (a) the circuit suppresses noise that may be present at the voltage supply output; in the fluidic case (b) the circuit suppresses pulsations from a pump.

for example, to abruptly stop flow in the microfluidic system, which may be necessary in order to perform a lengthy analysis on a sample. At the moment of the pump disconnection, the capacitor χ is still under the pressure difference ps = pρs /(ρ + ρs ). This pressure difference will continue to drive flow through the system as the membrane recovers to its equilibrium position. This situation is quite commonly encountered in electrical circuits when a capacitor discharges through a resistor. In that case the voltage in the capacitor decreases exponentially, with a time constant given by Rs C. The same thing happens in the fluidic circuit, and the pressure across the system will decrease following the same exponential law: ps (t) = p

t ρs − χρ e ρ + ρs

(3.69)

In other words, it will be impossible to stop flow instantaneously—there will necessarily be a time transient of order O(χρ). The more compliance the system has, the smoother (read: slower) this transition will be. If very fast flow switching is needed, the best bet is to work with hard materials that do not have a lot of compliance. This is an important criterion to keep in mind when designing a microfluidic system. 3.3.6.5 Viscous Heating Effects

Another well-known phenomenon from electricity is that of Joule heating. If a current I is injected through a resistor R, there will be an amount of heat

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generated per unit time that is given by I 2 R. As one might expect, the same is valid for fluidic circuits: fluid pushed at flow rate Φ through a capillary of hydrodynamic resistance ρ will generate, due to viscous heating, an amount of heat per unit time W given by the formula: W = Φ2 ρ

(3.70)

This can also be seen by using mechanical arguments: the energy wasted per unit time to push fluid through a capillary is equal to the force F that is applied multiplied by the average velocity of the fluid vavg : W = F vavg . The force, however, is equal to the pressure p times the area of the capillary S, but the average velocity multiplied by the area of the capillary is equal to the flow rate: Φ = vavg S, and hence we obtain that the amount of mechanical energy wasted per unit time (also known as the mechanical power) is equal to: W = pΦ. If we remember the definition of the hydrodynamic resistance: ρ = p/Φ, we recover (3.70) from purely mechanical considerations.

3.4 Surface Tension and Wetting As the dimensions of a system scale down, surface forces become dominant over body forces. Indeed, as the size l of the system is reduced, the ratio of the surface area to volume scales as l2 /l3 = 1/l (i.e., it diverges in the limit l → 0). Effects related to wetting and surface tension, for example, which at the macroscopic scale are often negligible, can become crucial at the microscale. Surface tension σ can be defined as the energy per unit area required to create an interface between two immiscible fluids (two liquids, or a liquid and a gas). The units are therefore J/m2 , which is equivalent to N/m (the more common unit). The energy of the interface is due to the fact that a molecule located at the surface of a fluid experiences fewer intermolecular bonds than the same molecule when located in the bulk of the fluid. Each intermolecular bond requires a certain amount of energy to be broken: for example, the strength of typical hydrogen bonds is on the order of 10 kJ/mol. Dividing this by Avogadro’s number, we obtain a hydrogen bond energy of approximately 1.7 × 10−20 J/molecule. Each molecule in bulk has several bonds, and we can assume that the molecules at the interface lose on average half their bonds—therefore, each molecule located at the interface accounts for approximately 8.5 × 10−21 J. If we assume an intermolecular distance of approximately 4 Å, we obtain an area per molecule of 1.6 × 10−19 m2 , which leads to an energy per unit area equal to σ ≈ 54 × 10−3 J/m2 (54 mN/m). It is interesting that this simple argument allows us to calculate a

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rather accurate value: most liquids have surface tensions in the range of 20 mN/m (for many aliphatic oils) to 80 mN/m (water). The tendency of a system to minimize its energy by reducing its interfacial area is responsible for many phenomena commonly encountered in microfluidics. The shape of droplets is such a phenomenon—the sphere is the geometrical shape with the smallest surface area for a given volume. A smaller area means a smaller surface energy, and is therefore the preferred configuration. Another such phenomenon is the instability that causes the breaking of elongated fluid cylinders (or jets) into droplets: the sum of the surfaces of the droplets is smaller than the surface of the cylinder; hence, the system migrates to the lower energy configuration. This is called the Rayleigh instability, and is quite often used in multiphase microfluidics to create monodisperse droplets using, for example, the flow-focusing technique [14]. 3.4.1 Surface Tension Forces There are forces and pressures originating from surface tension. Imagine that the free surface of a liquid film is pinned to a metal ring (a situation commonly encountered when blowing soap bubbles, for example). By increasing the radius of the metal ring (r → r + δr), we also increase the area of the film: 2πr 2 → 2π(r + δr)2 ≈ 2πr 2 + 4πrδr (the factor 2 in the area comes from the fact that a soap film has two interfaces). The change in total interfacial area is therefore 4πrδr and the additional surface energy is given by δE = σ · 4πrδr—the ring will therefore need to supply this amount of work: W = 2 · f · 2πrδr = δE (f being the radial force per unit length of the ring, per interface). From here we can infer: f = σ. In other words, each interface will act with a force per unit length equal to σ on the structure pinning it (Figure 3.12(a)); the surface tension forces are always located in the tangent plane of the fluid, and they act normal to the contact line. This is an important result, and it explains why the units most commonly used for surface energy are N/m rather than J/m2 . In the case of a spherical droplet of radius r, one can perform the imaginary experiment of cutting it in half: in this case, each half will be bounded by a circular line of radius r as well. The forces that act on this line are tangent to the fluid surface and nomal to the line, and they are therefore perpendicular to the imaginary cutting plane. The total force for each hemisphere amounts to 2πrσ. In the droplet, these forces can only be balanced by pressure forces, and therefore inside the droplet the pressure must be higher than outside by an amount: 2πrσ 2σ ∆p = = (3.71) πr 2 r

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Figure 3.12 (a) The two interfaces created by a soap film suspended by a metal ring act on the ring with a force f = σ per unit length and per interface (seen here in cross-section). (b) The triple contact line between two fluids (A and B) and a solid S, with the respective forces acting on the contact line.

This formula is known as the Laplace pressure. We notice that the pressure diverges for r → 0, which has some interesting consequences. In particular, very small gas bubbles tend to disappear very quickly in a liquid, as the enormous increase in pressure leads to rapid transport of the gas across the interface. It is worth mentioning here the Laplace pressure corresponding to nonspherical interfaces: ∆p = σ( r11 + r12 ), where r1 and r2 are the principal curvature radii at a point on the surface. 3.4.2 Wetting Considerations In the case where a film of fluid B is uniformly distributed over the surface of a solid S immersed in another fluid A, we obtain two planar interfaces: the B-S interface, with surface tension σBS , and the A-B interface (σAB ). Such a structure will be stable only if the energy corresponding to the layered structure is inferior to that of the A-S interface (σAS ). In other words, if σBS + σAB < σAS , we will have complete wetting of the solid S by the liquid B. Similarly, if σAS + σAB < σBS , we will have complete wetting by the liquid A. In case neither of these conditions is satisfied, the liquid layer will bead up forming droplets. Each droplet will contact the solid at a certain angle θ (called the wetting angle), as shown in Figure 3.12(b) (here we have represented the triple contact line between the two fluids, A and B, and the solid S; in the general case, the nonwetting condition becomes σAB > |σAS − σBS |). The wetting angle can be calculated quite easily, from the conditions of mechanical equilibrium of the contact line: σAB cos θ + σBS = σAS

(3.72)

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σAB sin θ = Felastic where Felastic is a vertical elastic force due to the deformation of the surface by the pulling of the surface tension. Typically, the deformations are extremely small and are completely negligible in most situations encountered in microfluidics. From (3.72) we obtain the value of the wetting angle: θ = arccos((σAS − σBS )/σAB ). This is only a theoretical value, however, as in practice there is a certain range within which the wetting angle can vary—this is the socalled phenomenon of contact angle hysteresis, and it is responsible among other things for droplet pinning on solid surfaces. The contact angle hysteresis originates in imperfections of the surface (roughness, contamination, defects) [15]. Wetting phenomena are extremely important in several situations in microfluidics such as the initial priming of microfluidic chips with very small channels and multiphase flows. Wetting properties of a surface can be tailored using a variety of techniques. Chemical patterning and nanoscale roughness, have both been used to create hydrophilic or hydrophobic surfaces, or portions of surfaces. Imagine that a microfluidic chip containing channels of radius r needs to be filled with a fluid. Consider a wetting angle θ between the fluid and the channel walls. In case θ < 90o , the channel will be filled by capillarity: indeed, the surface tension forces will tend to pull the fluid into the channel, greatly simplifying the priming job—this effect is called capillary wicking. If, on the other hand, θ > 90o , we have a problem: the capillary forces will push the fluid out of the channel, preventing priming. The pressure of the resulting meniscus (also called capillary pressure) will be given by: p = 2σ| cos θ|/r

(3.73)

This pressure needs to be overcome externally, and it can be considerable. As an example, let’s consider a PDMS chip with minimal channel radius of order 1 µm, which we want to fill with water. PDMS is hydrophobic in its virgin state, with a water contact angle θ ≈ 110o . The surface tension of water is approximately 80 mN/m; hence, the capillary pressure will be approximately 55 kPa, or about half a bar. In the case of multiphase flow, wetting conditions are obviously critical: if the surface of the channel does not have a strong preference for one fluid versus another, droplets will tend to stick to the walls and in the channel corners, thus disrupting the operation of the system. Several solutions are available to avoid such situations: different choice of materials, chemical treatment of the channel walls to make them preferentially wet by one of the

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two fluids, and use of surfactants that tend to reduce the surface tension (as we will see later). Most of the time a combination of these techniques is used. Another technique that is sometimes used to fill hydrophobic microfluidic channels with water is to prime them using a different liquid with smaller surface tension, which is miscible with water and wets the channel material. A common occurrence for PDMS chips is to use an alcohol (ethanol or isopropanol) for priming, and then replace it with water. Since the transition from alcohol to water is performed gradually without the presence of an interface, there will be no capillary pressure to overcome in this case. 3.4.3 Surfactants and Marangoni Stresses We are all familiar with the cleaning effects of soap: it allows water to remove grease and oils that tend to stick to surfaces, clothes, skin, and so forth. How exactly does soap act, to allow removal of such materials that are normally insoluble in water? The answer lies in the molecular structure of soap: the soap molecules are amphiphilic, which means that they have two parts—one that is soluble in polar solvents like water, the other that is soluble in nonpolar oils. When present in a single phase fluid, the soap molecules assemble into spherical micelles with the ends soluble in the fluid pointing out, and the insoluble ends directed toward the center of the micelle. This structure allows the soap molecules to reduce their total energy by reducing the contact of the insoluble (read: highly energetic) end with the fluid. In the presence of two phases such as oil and water, the soap molecules will migrate at the interface between the phases—each molecule in this case will have its hydrophilic end in the water phase and its hydrophobic end in the oil phase. This property of assembling at surfaces between fluids lead to the soap molecules being called surfactants. The main role of surfactants is to stabilize the interfaces between insoluble fluids and to reduce the interfacial tension (which, incidentally, is responsible for the cleaning effect). Quite naturally, the effective surface tension will depend on the type and the density of surfactant molecules present at the interface. Surfactants are used whenever stable droplets or bubbles need to be created: the surfactant molecules create a protective layer around the droplet that does not allow it to come into contact (and coalesce) with another droplet. This effect is applied in microfluidics when multiphase operation is required—surfactants (typically in the form of liquid soaps) are usually added to the water phase prior to operating the device. Surfactants also have the effect of rigidifying surfaces: while normally a liquid interface corresponds to a slip boundary condition in fluid mechanics, when the surface is charged with certain types of surfactants, it can

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Figure 3.13 The thermal gradient caused by a heater induces a surface tension gradient. The surface tension at the right end of the droplet is lower than that at the left end, causing a net force that pushes the droplet towards the heater.

become completely rigid—the appropriate boundary condition in this case is a no-slip one. As we have seen earlier when discussing the HadamardRybczynski relation (3.58), this phenomenon can create significant deviations of experimental results from theory. Such issues have been recognized a long time ago [16]; however, due to the huge variety of surfactants that can be present at the interface, it is difficult to arrive to exact theoretical results concerning the nature of the boundary conditions at fluidic interfaces in many real-world situations. Gradients in surface energy along the surface of a droplet or a bubble can lead to net stresses (or forces) on the interface or droplet. These are the so-called Marangoni stresses. There are different origins for the Marangoni stresses: the surface tension can vary along the interface either as a consequence of nonuniform surfactant coverage or as a consequence of thermal gradients (surface tension usually decreases with temperature, which explains why hot water washes better than cold water, and why oil spreads on top of hot water when cooking). Marangoni stresses can be useful in microfluidics and have been used as actuation mechanisms in a variety of situations [17]. As an example, we will consider the thermocapillary actuation of a droplet using the Marangoni stresses created by a temperature gradient. By patterning arrays of heaters under a channel containing a droplet (or by using heater arrays in an open configuration), one can control the temperature gradients externally, and thus guide the droplet [18, 19]. The idea, shown schematically in Figure 3.13, relies on the variation of surface tension with temperature. In the case of a nonwetting droplet (the most common situation in droplet microfluidics),

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heating one end of the droplet will lower the surface tension locally, and thus there will be a net force pushing the droplet toward the hot end. This can also act as a trapping mechanism: droplets, bubbles, and even cells can be trapped at hot spots. Unfortunately, most of the time effective thermocapillary trapping requires high temperatures that are incompatible with biological samples or even with some of the liquids used (causing them to boil). One can make a rough estimate of the force that can be generated using such a thermocapillary actuation method. The surface tension of water drops by approximately 0.2 mN/m per degree. If we consider a 10o difference in the temperature, this leads to a difference of δσ = 2 mN/m between the hot and cold ends of a droplet. If we consider the water droplet to be located in a hydrophobic capillary of radius r = 50 µm (wetting angle: 180o ), this leads to a net force on the particle equal to F = 2πrδσ ≈ 0.6µN. This is a very small force that corresponds to an equivalent pressure of 80 Pa pushing the droplet toward the hot end; nevertheless, in a properly designed geometry, this force can be sufficient to perform droplet actuation and manipulation actions.

3.5 Electrical Phenomena in Microfluidics As we have seen already, microscale physics deals with many phenomena that can be ignored at macroscopic dimensions. Electrical phenomena make no exception: dielectrophoresis, electro-osmosis, and electrocapillarity, for example, normally play no role in everyday life but can be harnessed and put to good use in microfluidics. 3.5.1 Electrophoresis Electrophoresis is the process of charged particle transport in a fluid under the effect of an electric field. We have already seen how to include the electrical forces in the momentum conservation equation for a fluid (Section 3.3.3). Now, instead of looking at the effect of the electric field on fluid flow, we will look at the effect on the motion of particles (or ions) through the fluid. Suppose we have a particle of radius r and carrying an electrical charge q suspended in a fluid of viscosity v. If we place the particle in an electrical field E, there will be a net force on the particle: Fel = qE. The particle will be accelerated until the electrical force will be balanced by the viscous drag force of the fluid, at which point the particle will have reached its terminal velocity v. We must have: qE = 6πηrv q from where we obtain the terminal velocity v = µE, with µ = 6πηr being called the mobility. This formula can be applied to charged colloidal particles

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as well as ions, with one caveat—r will not stand for the actual ionic radius, but for the hydrated radius (which includes a layer of polar water molecules accumulated around the ion). Since the terminal velocity will depend on the ratio between charge and hydrated ion radius (which depends on the type of ion), capillary electrophoresis is routinely used in analytical chemistry to perform separations based on ionic mobility. Electrostatic forces have also been used for droplet manipulation in microfluidic devices—a complete set of functions including drop generation, coalescence, and sorting has been demonstrated [20]. 3.5.2 Dielectrophoresis Dielectric particles that are not charged can be displaced through dielectric media using electric field gradients. The dielectric particle, when placed in an electric field E, will develop an electric dipole moment p = αE

(3.74)

where α is the polarizability of the particle. For a spherical dielectric particle of radius r, the polarizability α can be calculated using the Clausius-Mossotti formula: ǫ′ − ǫ 3 r (3.75) α = 4πǫ ′ ǫ + 2ǫ where ǫ′ is the dielectric constant of the particle and ǫ is the dielectric constant of the fluid. In an electric field E, it is possible to calculate the total force on a particle by integrating the electrostatic forces over the volume V of the particle: R F = V d3 rρ(r)E(r). We can now perform a Taylor expansion of the electric field about the center of the particle r0 and keep only the first-order terms (we assume that the particle size is small compared to the distances over which the electric field varies significantly so we can ignore higher-order terms in the expansion). We obtain then: F=

Z

V

d3 r(ρ(r)E(r0 ) + ρ(r)((r − r0 )∇)E(r0 ))

(3.76)

If weR recognize that Q = V d3 rρ(r) is the total charge of the particle, and that V d3 r(r − r0 )ρ(r) is the dipole moment p, we can rewrite (3.76) as: R

F = QE(r0 ) + (p∇)E(r0 )

(3.77)

The first term in this formula is the regular electrostatic force on the particle carrying total charge Q. If we assume the particle to be uncharged, then this

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term disappears. The second term is the force on the particle due to its induced dipole moment, when placed in an inhomogeneous field. To give an example, we can consider a simple field, oriented in the x direction, and that varies only depending on the x coordinate. In this case we can make use of (3.77) and (3.74) to obtain: ˆ α∂x Ex2 F=x (3.78) The particle will therefore be attracted (for α > 0) or repelled (α < 0), towards the maxima of the electric field. This effect is very important for microfluidic manipulations, since the forces can be switched on and off instantaneously (unlike, for example, forces arising from mechanical or thermal actuation, which suffer from a time lag due to the compliance of the system or from the thermal inertia). Dielectrophoretic actuation has been used, for example, to sort cells or droplets in real time and at frequencies up to 1.6 kHz [21]. 3.5.3 Electro-Osmotic Flow Electro-osmotic flow represents the movement of a globally neutral fluid relative to a charged solid surface under the influence of a DC external electric field [10]. It is a complex effect that can only be achieved successfully with certain fluids and channel materials and is quite sensitive to the chemical environment. Electro-osmotic pumping mechanisms are important in microfluidics because they generate a flat velocity profile that is not prone to Taylor dispersion, thus allowing a better resolution than pressure-driven systems in certain applications. The reason why this mechanism works has to do with the materials’ surface chemistries. Most materials, when placed in a fluid, develop a surface charge—glass, for example, typically acquires a negative surface charge when placed in an aqueous solution. Ions in the fluid migrate to the solid surface; some of these ions become attached to the channel walls, but the rest remain mobile, forming a very thin layer at the channel wall (called a Debye layer or a double layer). The ions in the double layer can move under the influence of an external electric field, and by viscous drag they pull the entire fluid column along. It is important to note that an electrochemical reaction needs to be possible at the electrodes, or else they would soon be screened themselves by Debye layers generated by migrating ions. A typical implementation of electro-osmotic pumping involves two liquid wells (one being the sample reservoir, the other being the water reservoir) containing electrodes, and connected by a channel. Flow is generated when a voltage is applied between the electrodes.

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AC electro-osmosis (ACEO) represents a similar technique of pumping fluids, which is becoming increasingly used in integrated microsystems. In ACEO systems, one wall of the microfluidic channel is patterned with interdigitated electrodes of alternating potential. The component of the electric field which is tangent to the electrode surface acts on the ions contained in the double layer, thus causing fluid flow [22]. As the polarity of the electrodes is reversed, so is the sign of the charges and the direction of the tangent electric field; the force on the double layer, therefore, always remains oriented in the same direction. For symmetrical electrodes, this force is oriented from the outer edge of the electrode towards the center, thus creating symmetrical vortices in the fluid. If the electrodes are of asymmetric design, the vortices above the narrower electrodes shrink and there is a net fluid flow in the channel above the electrodes. A limitation of ACEO is electrode erosion due to electrochemical reactions. This limitation can in theory be avoided by protecting the electrodes using a dielectric layer, which will require using higher AC voltages. It is important to mention that ACEO only works within an intermediate range of frequencies (usually around 500 kHz) which is related to the ion concentration and mobility. At lower frequencies, the ions in the solution have time to respond to the electric field and screen it, thus resulting in low ACEO fluid velocity. At higher frequencies, the net charge that develops in the double layer is small and so is the corresponding driving force on the fluid [23].

3.5.4 Electrowetting (Electrocapillarity) Electrowetting is another effect that is based on similar principles to electroosmosis, but which is used for manipulating droplets in two-phase flow. In the case of electrowetting, an electrical potential is responsible for disrupting the electrical double layer formed at the channel wall, which reduces the liquidsolid surface tension and, implicitly, the changes the contact angle (as we have seen in Section 3.4.2). A typical implementation involves an array of electrodes separated from the fluid by an insulating dielectric layer. Once an electrode is actuated, the drop meniscus closest to the electrode gets deformed creating a force on the droplet, which moves over the active electrode. Using this type of mechanisms it is possible to move, coalesce, mix, and break off droplets, thus opening a wide array of possibilities in digital microfluidics [24].

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3.6 Conclusions This chapter has covered the physics responsible for many of the phenomena encountered in microfluidic devices, and the material covered in this chapter will be referred to throughout the book. Since the phenomenology of microfluidic systems can be extremely rich, the number of covered topics is large. Due to space constraints, many topics could not be treated in depth. We tried to avoid tedious mathematics wherever possible, and to develop instead intuitive approaches. An effort has been made throughout the chapter, however, to provide enough information as to enable the reader to perform the basic calculations and back-of-the-envelope estimates required to design a microfluidic system correctly.

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References [1] Batchelor, G., An Introduction to Fluid Dynamics, Cambridge, UK: Cambridge University Press, 2000. [2] Lamb, H., Hydrodynamics, Cambridge, UK: Cambridge University Press, 1975. [3] Nguyen, N., and Wereley, S., Fundamentals and Applications of Microfluidics, Norwood MA: Artech House, 2002. [4] Bruus, H., Theoretical Microfluidics, Oxford, UK: Oxford University Press, 2008. [5] Landau, L., and Lifshitz, E., Fluid Mechanics, Oxford, UK: Pergamon Press, 1987. [6] Feynman, R., Leighton, R., and Sands, M., Lectures on Physics, Vol. 2, Menlo Park, CA: Addison-Wesley, 1964. [7] Myers, D., and Meyers, D., Surfaces, Interfaces, and Colloids: Principles and Applications, Weinheim, Germany: Wiley-VCH, 1999. [8] Einstein, A., “Investigations on the Theory of Brownian Motion,” Annalen der Physik; Translation: Dover, NY - 1956, Vol. 17, 1905, p. 549. [9] Pipe, C., Majmudar, T., and McKinley, G., “High Shear Rate Viscometry,” Rheologica Acta, Vol. 47, 2008, p. 621. [10] Stone, H., Stroock, A., and Ajdari, A., “Engineering Flows in Small Devices,” Annual Review of Fluid Mechanics, Vol. 36, 2004, p. 381. [11] Ajdari, A., Bontoux, N., and Stone, H., “Hydrodynamic Dispersion in Shallow Microchannels: The Effect of Cross-Sectional Shape,” Analytical Chemistry, Vol. 78, 2006, p. 387. [12] Moffatt, H., “Viscous and Resistive Eddies Near a Sharp Corner,” Journal of Fluid Mechanics, Vol. 18, 1964, p. 1. [13] Spice, http://embedded.eecs.berkeley.edu/pubs/downloads/ spice/index.htm. [14] Anna, S., Bontoux, N., and Stone, H., “Formation of Dispersions Using Flow Focusing in Microchannels,” Applied Physics Letters, Vol. 82, 2003, p. 364. [15] Quéré, D., “Wetting and Roughness,” Materials Research, Vol. 38, 2008, p. 71. [16] Scriven, L., and Sternling, C., “The Marangoni Effects,” Nature, Vol. 187, 1960, p. 186. [17] Darhuber, A., and Troian, S., “Principles of Microfluidic Actuation by Modulation of Surface Stresses,” Annual Review of Fluid Mechanics, Vol. 37, 2005, p. 325. [18] Sammarco, T., and Burns, M., “Thermocapillary Pumping of Discrete Drops in Microfabricated Analysis Devices,” Aiche Journal, Vol. 45, 1999, p. 350. [19] Darhuber, A., Valentino, J., Troian, S., and Wagner, S., “Thermocapillary Actuation of Droplets on Chemically Patterned Surfaces by Programmable Microheater Arrays,” Journal of Microelectromechanical Systems, Vol. 12, 2003, p. 873.

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[20] Link, D., Grasland-Mongrain, E., Duri, A., Sarrazin, F., Cheng, Z., Cristobal, G., Marquez, M., and Weitz, D., “Electric Control of Droplets in Microfluidic Devices,” Angewandte Chemie International Edition, Vol. 45, 2006, p. 2556. [21] Ahn, K., Kerbage, C., Hunt, T., Westervelt, R., Link, D., and Weitz, D., “Dielectrophoretic Manipulation of Drops for High-Speed Microfluidic Sorting Devices,” Applied Physics Letters, Vol. 88, 2006, p. 024104. [22] Ramos, A., Morgan, H., Green, N., and Castellanos, A., “Ac Electric-Field-Induced Fluid Flow in Microelectrodes,” Journal of Colloid and Interface Science, Vol. 217, 1999, p. 420. [23] Wong, P., Wang, T., Deval, J., and Ho, C., “Electrokinetics in Micro Devices for Biotechnology Applications,” IEEE/ASME Transactions on Mechatronics, Vol. 9, 2004, p. 366. [24] Pollack, M., Fair, R., and Shenderov, A., “Electrowetting-Based Actuation of Liquid Droplets for Microfluidic Applications,” Applied Physics Letters, Vol. 77, 2000, p. 1725.

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4.1 Introduction In the previous chapters we briefly reviewed the main physical concepts behind microfluidics, which provides us with tools to better understand the different phenomena present in fluidic systems at scales on the order of a few micrometers and smaller. We learned about different techniques that can be used to manufacture devices with fluidic conduits of that size and we implicitly understood the trade-offs between different technology paradigms. We learned about the different technologies that have been engineered to control where fluids go within microfluidic devices—how to design pumps, valves, mixers, and so forth, using different fabrication techniques. We also reviewed various sensing techniques that allow us to learn about what is going on within the few nanoliters of fluid present in our microfluidic system, and thus provide the functionality of the device. This chapter is where all the previous information converges. We are going to learn how to start from a concept and arrive at a functional sensor prototype. The journey will take us through every aspect of the design process: choice of technology, definition of the manufacturing process and sensor design and optimization (including computer-aided design (CAD), analytical estimates, and computational fluid dynamics (CFD) simulation). We will try to provide examples that capture the complete design workflow. Our choices of software programs is purely exemplary, is based on personal experience and preferences, and should not be considered as an endorsement in any way. Many commercial and free alternatives exist. The reader should always perform due diligence in selecting a certain equipment, manufacturer, or reseller. There are many other texts available that cover in more detail various aspects of the material presented in this chapter: MEMS and microfluidic manufacturing processes [1–7], design and photolithographic mask fabrication [8, 9], fluidic design, optimization, and computational fluid dynamics (using either commercial software or home made code) [10–13]. The reader is urged to consult these, and refer back to Chapters 1, 2, and 3 of this book for additional in-depth information about the topics covered.

4.2 Technology As we have seen in Chapter 1, several technologies can be used to manufacture microfluidic components. Four large categories stand out: • Soft lithography (PDMS); • Metallic fabrication;

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• Plastic fabrication; • Silicon/glass (MEMS) fabrication. In this chapter we will focus primarily on the two technologies with the greatest integration potential: soft lithography and silicon/glass fabrication. These technologies are becoming predominant in the microfluidic domain— they are the technologies most often applied to manufacturing individual fluid manipulation components (valves, pumps, and so forth), and for manufacturing the majority of available commercial microfluidic devices. Plastic manufacturing has become a viable alternative in recent years, but does not yet provide the miniaturization capabilities of either silicon/glass manufacturing or of soft lithography. The choice of one technology versus another is driven by a number of criteria that may weigh differently in each application. Some factors, like manufacturing costs, can often be a deciding factor in the choice of technology. Many variables may play a role in determining the most economical manufacturing alternative given a specific design. These may include, for example, the quantities involved (various technologies may scale differently with production volume), or the cost of external supporting hardware required for system deployment. We will not consider cost factors in our following analysis, for multiple reasons. First, cost calculations are very application-specific, and the cost estimation criteria cannot be extrapolated from one project to the next. Second, microtechnology evolves rather quickly, and certain methods that might have been prohibitive only months ago may quickly turn into viable manufacturing techniques. Finally, the current trend in microfluidic design seems to be towards increased added value per system, rather than towards cost reduction. This is natural, given that microfluidic technology is still in its infancy, and that there is an impressive number of markets where the benefits of microfluidics have not yet penetrated: many efforts are concentrated at enabling new applications rather than producing low-cost devices for existing markets. There are exceptions to this, of course, the most notable one being the ink-jet industry, where cost considerations can be important in the manufacturing of print heads. 4.2.1 Functional Description A microfluidic project starts with a description of the required device functionality. Often, the microfluidic project involves a miniaturization of an existing laboratory measurement or process. In this case, the different sample manipulations that are required to perform the measurement are imposed by the laboratory protocol, and the microfluidic device must perform the right

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sequence of steps at the right moments. Nevertheless, creating the microfluidic counterpart of a laboratory experiment is not a trivial task, since the behavior of the fluids can be very different at large scales from that in a microfluidic device. Often, this difference can have undesirable outcomes. Mixing is a good example: while large-scale systems benefit from the chaotic mixing induced by turbulence, microfluidic devices usually operate at very low Reynolds numbers where flows are laminar and mixing is difficult to achieve. Therefore, scale-specific mixing mechanisms must be implemented to achieve reliable mixing of components. In other cases, device miniaturization can enable creative ways to achieve the required sample manipulations. A good example of how a microfluidic implementation can be different from its laboratory counterpart is given by thermal cycling as used in the polymerase chain reaction (PCR). In a laboratory PCR procedure, thermal cycling is typically achieved by changing the temperature of the entire sample using a programmable hot plate or equivalent—the preprogrammed temperatures for the different steps in the PCR procedure are therefore imposed on the entire sample. In a microfluidic device, on the other hand, one can impose a temperature gradient along one direction of a sample and create a serpentine channel that guides the fluid successively through high and low temperature regions (Figure 4.1). The temperature cycling in this case is not achieved by successively heating and cooling the entire sample, but by circulating the fluid between hot and cold regions of the device. This procedure would be difficult to achieve in a laboratory-scale experiment, but is relatively simple to implement in a microfluidic geometry. The main interest of microfluidic systems does not lie, however, in merely miniaturizing existing laboratory procedures, but rather in taking advantage of the scale reduction to enable novel applications, or to perform a chain of specific manipulations that are difficult if not impossible to achieve on a large scale. In this case, the scientist must divide the process into individual steps that can be achieved using standard building blocks such as those described in earlier chapters of this book, or must design his or her own application-specific building blocks. 4.2.2 Integration: Monolithic Versus Hybrid From earlier chapters we learned that most microfluidic functional building blocks can be implemented in more than one way. It is therefore important to keep the choice of technology and manufacturing process as flexible as possible. Sometimes, all the required functional blocks can be implemented

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Figure 4.1 Polymerase chain reaction can be achieved in a microfluidic chip using a completely different procedure from the standard laboratory one: thermal cycling can be imposed by circulating the DNA sample through a serpentine channel in a thermal gradient rather than by heating and cooling the entire sample.

monolithically within a single technology (highest degree of integration); other times several technologies need to be combined in a hybrid system (highest versatility). The decision to proceed one way or another is critical in the early stages of system design, as it will define the limitations and the accuracy of the system. Monolithic integration is desirable whenever possible. Having all the functional building blocks and connecting channels integrated on a single chip substrate results in many advantages, which can be critical in certain applications. The dead volumes that typically appear in fluidic interconnects are eliminated, resulting in rapid response times, reduced sample contamination, and a decrease in the sample and reagent volumes used. By reducing the number of fluidic interconnects, a higher accuracy can also be achieved in many measurements. The physical size of the final system is significantly reduced, due, on one hand, to the reduction in the number of required fluidic interconnections (which consume a lot of chip space), and, on the other hand, to the reduction of wasted chip space (due to the safety zone around the perimeter of a device, where no channels or devices are implemented). The manufacturing process can be streamlined: a single process flowchart is required for manufacturing the device; once the fabrication process is well defined, this can result in fewer manufacturing errors and in significantly reduced cost per device. An additional advantage comes form the fact that the fluid comes in contact with a smaller number of

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Table 4.1 Trade-Offs Between Monolithic and Hybird Integration

+

–

Monolithic Integration

Hybrid Integration

Dead volume reduction Accurate measurements Rapid response time Size reduction Streamlined manufacturing Cost advantage Trade-offs required Requires strict process control Lack of versatility Extensive simulations required

Versatility Best solution for each component Individual component optimization Flexible manufacturing Important dead volumes Slower and less accurate response Poorer chemical compatibility Large system footprint Additional manufacturing steps

materials, which reduces possible chemical compatibility issues. There are certain disadvantages of monolithic integration as well. By being forced to use a single technology for manufacturing all the building blocks, the designer may be tempted to make certain trade-offs and to use less-than-optimal solutions just because they fall within the capabilities of the technology. The manufacturing process is complex and the process control needs to be very strict—a single manufacturing error can compromise the entire system, leading to a high price being paid for a useless device. The design is much less flexible, and changing one parameter of one component of the system will require remanufacturing the entire system. Careful systemlevel simulations are typically required before launching a manufacturing run to ensure that all components of the system work well together. In hybrid integration approaches, by contrast, different functional blocks are manufactured using specific technologies and then are integrated within a complete system by using different types of fluidic and electrical interconnects. Hybrid integration is very versatile, as it allows each system component to be easily replaced in case it does not perform well. Individual building blocks can be optimized independently within the bounds of their respective technologies, and then interconnected to obtain a complete functional system. Different configurations can be attempted using the same building blocks, which can result in rapid optimization. Manufacturing

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process control is less stringent, as individual components can be easily replaced if found defective. The trade-offs of hybrid integration are important, and often outweigh the benefits. Due to the increased dead and swept volumes found in the fluidic interconnects, hybrid systems tend to be less responsive and less accurate. Being manufactured of different materials, microfluidic components may have conflicting chemical compatibilities (e.g., while plastics have excellent acid but poor solvent resistance, silicon, glass, and metallic materials have good solvent resistance but relatively poor acid resistance). A hybrid system manufactured employing both classes of materials will have poorer chemical compatibility than either one of them. The physical size of a hybrid system will necessarily be larger than the monolithic counterpart, often significantly so: fluidic interconnects, in particular, are very space-hungry. The manufacturing cost will typically be higher than for a monolithic system, as each component requires additional manufacturing steps when manufactured independently. In addition, the system requires a microassembly step that is absent in the case of monolithic integration: all the components need to be connected together to form a functional system. It is tempting to make the analogy between electronic and fluidic circuits. Hybrid integration seems to be similar to using a breadboard to manufacture an electronic circuit, whereas monolithic integration seems to be closer to designing an integrated circuit. On a breadboard one can perform a lot of troubleshooting and easily change components until the desired circuit function is achieved. Likewise, in hybrid fluidic systems, the fluidic building blocks can be changed and tweaked in an attempt to optimize the system. The design of an integrated circuit on the other hand is much less flexible, and often extensive simulations are required before a certain design is validated and manufactured. The same can be said of monolithic fluidic systems, where simulations (both order-of-magnitude and complete CFD modeling) are critical in system design. The relative advantages and disadvantages of the two approaches to fluidic integration are summarized in Table 4.1. There is a major difference, however, between electronic and fluidic systems: while in many electronic applications the size of connection wires between the components on a breadboards does not have a direct impact on circuit operation, in fluidic systems the connections are absolutely critical. A high dead volume in a fluidic connection can increase the response time of a fluidic system from milliseconds to tens of seconds. Using a flexible tubing can have a similar impact due to the pressure compliance of the tube walls. There is no such thing as a universal breadboard for fluidic circuits— this is a major obstacle for hybrid integration, and the solution is far from

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Figure 4.2 An example of channel plate use for providing multiple microfluidic building blocks with fluidic and electrical connectivity. The sealing of the fluidic ports can be achieved using O-rings, gaskets, or adhesive connections, whereas the electrical connectivity can be realized using wire bonding techniques or technologies borrowed from flip-chip electronics.

trivial. Usually it involves designing channel plates that are applicationspecific and manufactured using noncompliant materials. The channel plate provides fluidic connectivity between the different fluidic blocks, which are either adhesively attached to the plate and sealed or use O-ring connections or gaskets. The channel plate can also be provided with metallic traces, similar to a printed circuit board, which provide electrical connectivity and power to actuate valves and pumps, or to perform certain measurements. The electrical connections to the different microfluidic blocks can be achieved using wire bonding, or flip-chip techniques (Figure 4.2). The channel plate may even incorporate CMOS electronics, which can provide the system with on-board intelligence such as the ability to digitize measurements and communicate the results to other devices. Hybrid integration can also be achieved using commercial fluidic ports adhesively attached to the devices. In this case the interconnections between different functional blocks are made using external capillary tubing. This technique is versatile, particularly in the prototyping stage; however, it is time-consuming to manufacture (requiring manual port positioning and gluing), and wastes a lot of device space. The natural progression in designing an integrated microfluidic system is to start with designing and understanding individual components. Once these are well characterized, they can be integrated in a hybrid system, initially using commercial fluidic ports, and later using dedicated channel plates. Once the system design is well defined, an attempt to monolithic integration can be made. Usually, several manufacturing runs are required

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to properly tweak all the manufacturing processes and to get the individual functional components to work well in a complete system. This is a timeconsuming process, and requires good process control and clean room expertise. Certain designs cannot be realized in a monolithic system, either due to technology incompatibilities between individual components or to prohibitive manufacturing complexity. The decision whether to make a monolithic or a hybrid design needs to be made on a case-by-case basis. 4.2.3 Material Choices Every application requires materials that are chemically and biologically compatible with the fluids being used, which can withstand the physical parameters of the process (most notably, temperature and pressure), and which are compatible with the type of measurement involved (an optical measurements will require transparent materials, an electrochemical measurement will require conductive electrodes, and so forth). The choice of materials is an important decision, which is made early on in the design process, and is intimately related to the choice of manufacturing technology. Physical, biological, and chemical requirements are imposed by the application at hand. Among the physical parameters of interest are the temperature and pressure at which the devices must operate. High pressures require relatively rigid materials, with high yield stress—metallic and silicon/glass manufacturing might be the ideal choice for demanding environments where high temperatures and pressures are present, such as the chemical and petroleum industries. Biocompatibility is required in applications where live matter needs to analyzed (such as cells and bacteria), or whenever an implantable system is designed. In this case, the best choices may be inert plastics, elastomers, silicon, and glass. Chemical compatibility needs to be assured between the materials used in manufacturing and the fluids of interest. In the case where organic solvents are used, the ideal materials are silicon, glass, and metals. On the other hand, acidic environments may corrode or etch silicon oxide, glass, and metals; the ideal choice in that case might be the use of an inert plastic such as PEEK or PTFE. The properties of the manufacturing material may be important for the actual device function; for example, devices using resonant principles require relatively stiff materials with excellent elastic properties, so monocrystalline silicon is the ideal choice in this case. Similarly, the need for optical clarity may be a requirement in applications where optical investigations are performed (fluorescence and absorbance measurements, or general microscopy). In this case the material of choice is glass, along with silicone elastomers (PDMS) and clear plastics (acrylates, cyclic olephin copolymers, and so forth).

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4.2.4 In-House Versus Outsourced Fabrication The technology of choice for manufacturing a particular microfluidic system is intimately related to the choice of materials. Plastic fabrication may be best approached by precision injection molding or conventional machining; bonding in this case can be achieved by laser welding, adhesive, and solventassisted bonding. Glass and ceramic fabrication are best served by MEMS-like deposition and wet etch processes, high-speed machining, or sandblasting; bonding can be achieved by adhesive methods or by high-temperature fusion and anodic bonding techniques. Metallic fabrication requires machining or LIGA fabrication processes; brazing and diffusion bonding processes can be used for sealing off the system. Silicon manufacturing can use the full array of MEMS deposition, dry and wet etch processes; appropriate bonding processes include anodic, fusion, and eutactic bonding. Elastomers are best manufactured by hot curing and molding off SU-8 patterned master templates. Plasma-activation and variations in the initiator concentration can be used to bond two elastomeric substrates or an elastomer to a glass slide. Manufacturing an integrated microfluidic system requires access to at least one and, most likely, to multiple such manufacturing technologies. In the case of hybrid integration, in particular, individual components are likely to be manufactured using specific technologies. These technologies may be available in-house; if not, then outsourcing will be required for at least some fabrication steps. Even in the case of monolithic integration, where a single type of technology is used for manufacturing the entire system, certain process steps may not be available in the local clean room or manufacturing facility, thus requiring some outsourcing. Any clean room or microfabrication facility maintains a list of reliable partners that offer technologies not available inhouse, so that a complete manufacturing solution can be provided. Ideally, however, the number of outsourced fabrications steps should be kept to a minimum, in order to reduce the risk of error due to miscommunication of the specifications or to mishandling during shipping and manufacturing. In recent years, outsourcing of complete MEMS and microfluidic project fabrication has become increasingly common. A number of university and commercial foundries offer fabrication services, either using internal staff or allowing external users to access fabrication equipment (Figure 4.3). This situation benefits both the manufacturing facilities, which generate additional budgets for purchasing and maintaining equipment, and the external users, which have access to very expensive equipment without capital expenditures and maintenance fees. Such outsourced manufacturing services are particularly important for small start-up companies built around a good idea with market potential, and who are willing to pay for manufacturing

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Figure 4.3 The Cornell NanoScale Science and Technology Facility at Cornell University (Ithaca, New York) is an example of outstanding microfabrication facility that supports a broad range of microscale and nanoscale science and technology projects, with over 700 users per year, of which more than half are external to Cornell University. (Reproduced with permission from [14].)

prototypes and validating the idea without the large investments required for building and equipping a clean room. Some foundries provide standard fabrication processes and specify complete design rules to make sure the fabrication runs smoothly. Other foundries allow the users to tweak the parameters of the fabrication process. Depending on the foundry type, more or less flexibility may exist in designing a custom process. Typically, university clean rooms tend to be more open and allow wider access to the machines and to process parameters. Commercial foundries are less likely to allow modifications of standard processes that have been optimized for their equipment. The choice whether to invest in internal fabrication facilities or to outsource the fabrication depends on several factors. A first factor is how mature the design and the technology are—in case extensive process development is envisioned, it may be wise to first use an external partner for process and prototype development and to invest in critical manufacturing equipment only once the fabrication process is established and the prototypes are validated. A second factor is the available financial and human resources. A careful economic and risk analysis should be made prior to investing in a clean room and in expensive manufacturing equipment and trained personnel. Running a clean room is expensive and can quickly deplete the budget of a small company if it does not generate revenue. Finally, there is a third factor, which is the strategic plan. Is there a risk of exposing important know-how or other type of intellectual property by using a third-party manufacturing facility? Is it important to maintain a degree of manufacturing flexibility that cannot be achieved if the manufacturing is outsourced? Would an internal manufacturing line enable other potential sources of revenue that might

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offset the costs, for example, by offering foundry services? Answers to such questions need to be given on a case-by-case basis. In case the decision is to outsource the fabrication, several possibilities open up. One can either use commercial foundries and manufacturing shops or university partners. Generally, if the manufacturing process is standard and is being offered by a commercial foundry, then that is the most straightforward way to proceed. If the design guidelines are followed, then the foundry is responsible for providing the expected prototypes; the only risk assumed in this case is that the design itself may be flawed, in which case one or several additional iterations may be required. If, on the other hand, a lot of process development is required, then contracting a university clean room may be the best solution. Most of the time university clean rooms only provide personnel for maintaining the machines and for training other users but not for performing the actual manufacturing. The equipment needs to be operated by trained personnel of the contracting company or by the university staff (typically graduate students, postdocs, and technicians). Some university foundries offer fabrication services that are similar to those of commercial counterparts and provide trained technicians and engineers who perform the actual manufacturing. In this case, the contracting company may benefit from a good balance between flexibility and fabrication know-how provided by the clean room staff. 4.2.5 Fabrication Process: Definition and Optimization The design of the sensor is intimately related to the manufacturing process. Sometimes, a fabrication process needs to be developed to account for the details of the design; other times the design is tributary to a specific fabrication process that needs to be followed. Most of the time there is an interplay between these two dimensions of sensor development, and one needs to navigate back and forth between design and fabrication process to determine the details of either one. The manufacturing process needs to be established as soon as the choice of materials has been made. A detailed design is not required at this point; however, the general lines of the design need to be understood before the proper fabrication steps for manufacturing it can be determined. A diagram which describes a relevant cross-section through the device is usually sufficient at this stage. The diagram should, however, display the full complexity of the sensor structure, with all required fabrication layers. It is a good idea to discuss the sensor structure and the manufacturing process with the clean room or the manufacturing shop technical staff. Such people accumulate an enormous amount of information by working on hundreds of

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Figure 4.4 Example of a published micropump fabricated using multilayer soft lithography, which is the subject of one of the design examples described in this chapter. Left: The micropump integrated with valves and detection regions to create a sorting device (scale bar: 1 mm). Right: A detail of a micropump, showing the three actuation channels and the fluidic channel separated by a thin deformable PDMS membrane (scale bar: 200 µm). (Reproduced from [15] and [16], respectively.)

different projects, and can usually point out errors in the design or difficulties in the manufacturing process that could be avoided by a small change in design or technology. Always talk to the shop personnel early on in the design process, and then again to finalize the manufacturing plan. To demonstrate how a microfluidic device might evolve from the idea stage to a complete system, we will consider two examples and follow the different phases of the design process. The two examples that will be considered are a micropump based on multilayer soft lithograhy (using PDMS as the elastomer) and an emulsion generator manufactured in hard materials (silicon/glass). 4.2.5.1 Micropump Fabrication Process

We will outline a possible fabrication process for a micropump based on multilayer soft lithography [16,17], a technology capable of creating complex three-dimensional structures that can then be used as fluidic manipulation components (valves, peristaltic pumps, and so forth). The technology relies on low Young’s modulus materials such as silicone elastomers (PDMS) that are cast in thin deformable membranes sandwiched between more rigid

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Figure 4.5 A relevant cross-section through a microfluidic pump device shows the access holes for both pneumatic actuation and for fluidic connections, the thin PDMS actuation membrane and the two thick PDMS substrates on either side, the right-angle connection between the actuation and the fluidic channel, and the rounded corners of the fluidic channel. All these features need to be defined and accounted for in the fabrication process.

substrates to create pneumatically actuated valves and pumps. We will attempt to reproduce a design similar to that published by Quake et al. [15, 16], where a sequence of three valves are actuated peristaltically to create flow (Figure 4.4). A relevant cross-section through an exemplary device is shown in Figure 4.5. The device includes two channels crossing at right angles and separated by a thin membrane. Pressure is applied in the top channel and the membrane deflects downwards, thus closing the bottom channel. Depending on the actuation pressure in the top channel, the bottom channel may close partially or fully. It is important to note that the cross-section of the channels plays an important role: square cross-section channels cannot fully close even at very high actuation pressures (the membrane is not capable of deflecting enough to completely block corner flow), whereas rounded channels provide complete sealing at accessible actuation pressures [16]. Once the required features have been defined, the manufacturing process can be outlined. An initial manufacturing chart will define the main components to be manufactured and the general processes that will be used. In the case of the multilayer lithography process we are attempting to use, we need to define three different components: the top layer (TOP), which needs to include the pneumatic channels and access holes; the membrane layer (MEMB), which includes the fluidic channels; and the bottom layer (BOT), which includes the fluidic access holes. Each of these components will be made by molding a PDMS elastomer on a master mold, peeling it

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Table 4.2 An Initial Fabrication Chart for a Multilayer Microfluidic Device Using Soft Lithography

Initial process chart—Micropump

Step

Description

Comments

1 1.1 2 2.1 2.2 3 3.1 4 5 6 7 7.1 7.2 7.3 7.4

TOP layer master mold Photolithography—TOP-mask MEMBrane layer master mold Photolithography—MEMB-mask Curing BOTtom layer master mold Photolithography—BOT-mask TOP layer molding MEMB layer molding BOT layer molding System assembly TOP hole punching TOP+MEMB alignment/boding BOT hole punching TOP+MEMB+BOT alignment/bonding

Si wafer substrate Actuation channels Si wafer substrate Fluidic channels Photoresist reflow Si wafer substrate Fluidic access holes Silicone elastomer Silicone elastomer Silicone elastomer Alignment and bonding Pneumatic access holes Fluidic access holes

off, and aligning and bonding the different components to form the complete system. The fabrication will therefore include both manufacturing the master mold and fabricating the PDMS components. The initial fabrication chart needs to clearly identify every photolithography step, as these steps will require separate layers in the CAD design software that will lead to separate lithography masks. An initial chart for the multilayer lithography process may look like Table 4.2. Individual steps from Table 4.2 will in turn be developed further, until the fabrication process is broken down into individual steps that are clearly defined. The type of information that might be included for a photolithography step might be the type of lithography mask used (polarity and material; quartz masks have lower UV absorbance than polymer or transparency masks, and therefore require shorter exposure times to achieve the same radiation dose), the type of photoresist (positive/negative, manufacturer and type), the spin and ramp speeds used in spin coating (optionally, the thickness achieved after

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spin coating as measured using ellipsometry or profilometry techniques), the substrate pretreatment (was an adhesion promoter used?), any baking steps that need to be performed prior to or after exposure (temperature and duration, and type of baking; oven or hot plate), the alignment procedure performed (frontside or backside alignment? which alignment marks were used?), the type of exposure (contact or proximity), the total exposure time and intensity (alternatively, the dose may be specified in J/cm2 ), the type of developer used and the total required development time, any additional curing steps required for achieving the desired photoresist mechanical properties or shape (a reflow step may be needed to obtain channels with round corners), and any washing/drying procedures. For the molding steps, one needs to specify the master mold used, the type of elastomer, and the mixing ratio between the base product and the initiator (RTV silicones used in microfluidics usually are used as a twocomponent mixture), the process for dispensing the silicone onto the mold (total thickness, if poured; ramp, spin speed, and total spinning time, as well as the resulting thickness if a spin coater is used), the type and duration of the degassing step (required to get rid of air trapped within the details of the microstructure, or during mixing within the elastomer itself), any surface treatments that have been made to avoid adhesion to the mold (if a vapor treatment, then the type of product and total exposure time to the vapors need to be specified), and the time and temperature of the different curing steps. A complete process chart for such a manufacturing process may require close to a hundred individual steps, each with detailed specifications. Most of the time these steps are not explicitly written down; they are usually implied, and prior experience with the respective processes and products will be applied. In case a process needs to be outsourced to a fabrication facility, where a technician who may not be familiar with the process needs to perform the manufacturing steps, this type of detailed process chart is very important. It represents a recipe, and should allow for a trained person to understand the process in all details and to follow it through. This process chart contains the complete know-how about the fabrication of the device, and therefore is very valuable and is usually kept secret. Even scientific publications and patents typically only outline certain critical steps of the manufacturing process. 4.2.5.2

Emulsion Generator Fabrication Process

A second example of device is an emulsion generator fabricated in siliconglass technology, which uses a concept developed by Sugiura et al. [18]: plane pancake-like droplets that are constrained within a narrow slit develop an instability when they reach a step edge—the fluid forms a spherical droplet

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Figure 4.6 Left: The design used by Sugiura for high-throughput emulsification consists of a micromachined silicon chip in contact with a glass slide (shown separated for clarity). Right: A schematic of the monodisperse droplet formation process, showing the formation of the pancake-like droplet (a–c) constrained by a shallow terrace, and the development of the instability at the step edge (d–f) that results in the periodic formation of droplets. (Reproduced from [18]. Copyright 2001 American Chemical Society.)

that has smaller interfacial energy and splits off from the pancake-like droplet. The process then repeats periodically, resulting in the formation of a stream of monodisperse droplets. Multiple channels can produce droplets in parallel, which has the pontential for high-throughput emulsification. The original design of Sugiura et al. was manufactured in silicon, using photolithography and orientation-specific etching, and the silicon chip was pushed against a glass slide. Figure 4.6 shows the resulting geometry. The emulsion was collected in the macroscopic volume around the chip; it was not possible to capture the emulsion in a sealed microchannel due to the lack of bonding between the glass slide and the silicon chip. We will attempt to design a similar device, which uses multilayer photolithography and deep reactive ion etching (DRIE) of silicon as the manufacturing technology. We will, however, try to build a fully contained microfluidic chip, which uses two inlets (for the continuous and the dispersed phases) and collects the emulsion in a microchannel connected to an outlet from where the emulsion can be guided to a different microfluidic device if necessary. Alternatively, the emulsion generator could become a building block in a more complex microfluidic device. The emulsification channels will be etched in silicon using two etch depths, and access holes for the

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Figure 4.7 Left: A schematic of the emulsion generator device described in the text. The dashed white line represents a portion of the cross-section shown at right. Right: Schematic cross-section through the device, showing the dispersed phase channel and its inlet hole, the channel and terrace, and the outlet channel with the outlet hole (not to scale).

fluid inlets and the emulsion outlets will be machined through the silicon; a glass layer will seal off the device, allowing optical microscopy investigations of the emulsification process. All the micromachining will be performed on silicon, the glass being used only in the final anodic bonding step. The device geometry will mimic that of Sugiura, with a terrace etch depth of 2.5 µm and width of 12 µm, a channel width of 4 µm, and fluidic channels (inlets and outlet) 100 µm deep and 500 µm wide. The details of the geometry are shown in Figure 4.7 (left), along with a relevant cross-section through the device (right). As it is apparent from Figure 4.7, the device required three different lithographic processes, corresponding to the three etch depths: the inlet and outlet holes (through-wafer backside etch, mask name: HOLES), the channel and terrace (2.5 µm frontside etch, mask name: TERRACE) and the fluidic channels (100 µm frontside etch, mask name: CHANNEL). DRIE etching requires an etch mask, which can be made from aluminum for large etch depths, and from photoresist for small to moderate etches (typically below 200 µm). For our device, the fluidic channels will be etched to a depth of 97.5 µm using combination of photoresist and aluminum masks (aligned using frontside alignment (FSA)), then the photoresist mask will be removed by a solvent bath, thus exposing the channel and terrace regions to etching. The etch will be continued for an additional 2.5 µm using only the aluminum mask. The access holes will be etched from the back side using another aluminum mask (and backside alignment (BSA)). This procedure is shown schematically in Figure 4.8, and the corresponding fabrication process is summarized in Table 4.3. In a complete process chart, the fabrication steps should be detailed

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1 2 3 4 Figure 4.8 The silicon etch procedure for realizing the geometry of the emulsion generator. 1: The aluminum and photoresist masks are manufactured using photolithography and wet etch techniques. 2: The fluidic channels are etched 97.5 µm. 3: The photoresist mask is removed, and etching is continued for another 2.5 µm—the total etch depth for the fluidic channels will be 100 µm, and for the terrace—2.5 µm. 4: The wafer is etched from the backside through an aluminum mask to open the inlet and outlet holes.

completely. Some of the details that need to be included have been specified during the previous section. Additional details that need to be specified are listed next. A deposition step should list the type of deposition (sputtering or evaporation), the substrate temperature (which accounts for the film structure), and the total deposition thickness. An etch bath should describe the chemical composition and temperature of the bath as well as the etch time. A DRIE step should describe the type of process used (Bosch, cryogenic, and so forth), the plasma parameters, the substrate temperature, and the total etch time and resulting etch depth. An anodic bonding step is specified by the temperature of the substrates, the applied voltage, and the total bonding time; the total charge transferred during bonding can also be recorded. A very important consideration in designing a fabrication process is the order of the fabrication steps. While seen from the design perspective, there may appear to be no connection between two fabrication steps, but in practice they may need to be accomplished in a certain order. A simple example may consist of a wafer that needs to have some regions covered by a metal layer, and other (separate) regions covered by a silicon nitride layer. While the two steps could, in principle, be performed in any order, in reality they are not interchangeable—LPCVD nitride deposition is typically performed at very high temperatures (between 750o C and 1,000o C) that are close to or above the melting temperature of many metals, particularly aluminum and gold, which are predominantly used. Even if a high melting point metal like platinum

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Table 4.3 Initial Fabrication Chart for an Emulsion Generator Using Silicon/Glass Technology

Initial Process Chart—Emulsion Generator

Step

Description

Comments

1 1.1 1.2 2 2.1 2.2 2.3 2.4 2.5 2.6 2.7 2.8 3 3.1 3.2 3.3 3.4 3.5 4 4.1 4.2 5

Substrate preparation Preparation for processing Al deposition front/back Frontside fabrication Photolith. frontside CHANNEL mask Al etch Photoresist removal Photolith. frontside TERRACE mask DRIE etching Photoresist removal DRIE etching Al etch mask removal Back side fabrication Photolith. backside HOLES mask Al etch Photoresist removal DRIE etching Al etch mask removal Bonding Preparation for bonding Anodic bonding Dicing

Si wafer substrate 450µm Wafer cleaning Evaporation/Sputtering Si wafer substrate + Al layer Wet etching Solvent bath Thick resist; FSA 97.5µm Solvent bath 2.5µm Wet etching Si wafer substrate + Al layer BSA Wet etching Solvent bath 350µm (=450-97.5-2.5) Wet etching Silicon and glass Wafer cleaning Diamond saw

is used, the prolonged exposure to high temperatures will induce significant changes in the properties of the metal, resulting from rapid grain growth. In addition, clean rooms usually choose to maintain LPCVD furnaces free from metallic contaminants, thus assuring that the properties of the silicon nitride film are highly reproducible. Therefore, silicon nitride deposition is typically performed before most other processing steps. A notable exception is wafer oxidation, which can be performed both before (standard procedure) or after nitride deposition and patterning (in a local oxydation (LOCOS) process).

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4.3 System Design and Optimization The detailed system design can be accomplished only after the fabrication process has been correctly specified. A minimum set of process specifications need to include the types of materials used (both for the substrates and for the deposited films), the details of the photolithography steps involved, and the details of the associated processes. Typically an initial process chart like those outlined in Tables 4.2 and 4.3 needs to be specified. Once the process is clearly understood, one can proceed to the next phase: that of system layout and design. This is typically achieved using CAD software packages. For most microfluidic applications, two-dimensional layouts are required—after all, the majority of microfluidic systems are manufactured using at least some photolithography steps, which involve the design of a planar lithographic mask. We will start this section by explaining how a CAD software can be used to create a lithographic mask, and what kinds of features need to be included in the mask to allow alignment between the different photolithography layers. We will then learn how to optimize the layout and sizing of fluidic channels to achieve the desired microfluidic system behavior. To this end, we will apply order-of-magnitude estimates resulting from approximate physical models in the first place, to determine roughly the system behavior. Then we will apply some component-based simulation concepts, which allow us to replace fluidic channels by their electrical equivalents and model the microfluidic system as an electrical circuit (for which numerous simulation tools exist). Finally, we will perform detailed analyses of the fluidic behavior of critical components using computational fluid dynamics (CFD) software. These three levels of refinement should be sufficient in the majority of cases to determine with certain accuracy the fluidic behavior of the system. There will be situations, however, where the system is too complex to be modeled with any of these techniques. In those cases, the only solution is to experiment. Several prototype designs involving variations in the layout will need to be created, and the experimental results analyzed to select the best design that will be used in production. 4.3.1 Mask Design The lithography masks represent the blueprint of the microfluidic system. The features drawn on the masks (the “digitized data” as the mask manufacturing jargon calls it) will turn into metallic traces, channels, or through-holes, depending on the process used in conjunction with that photolithography step. In addition to the device features, a mask needs to contain registration or

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alignment marks. These are patterns that appear on the first lithographic mask used in the process and on each successive mask, to allow accurate registration between successive photolithography steps. Masks come in two possible polarities: positive (in which case the digitized data appears dark on a transparent background), and negative (digitized data transparent of a dark background). Depending on the type of photoresist used and on the type of process, a photolithography step may involve either a positive or a negative mask. The polarity of each mask will be specified when the masks are ordered from the mask manufacturer. Lithography masks can be printed on different substrates. The most accurate (and most expensive) masks are printed on a chrome layer that is supported by a quartz glass plate. These masks can reach submicron resolutions (features down to 500 nm can be created using standard lithography processes and machines); however, cost increases exponentially when reducing the minimum feature size. Masks can also be printed using chrome on a regular glass substrate (borosilicate or soda-lime glass), or using high-resolution photographic emulsion on glass or polymer film. The cost of a 5-inch chrome mask with 2 µm minimum feature size printed on glass is approximately $500. Polymer masks are the least expensive, the cost of a 5-inch polymer mask with 15 µm minimum feature size reaching as low as $50. Polymer masks, however, provide the poorest resolution and registration accuracy. Minimum feature sizes that can be printed on polymer masks are on the order of 5 µm, and registration errors of several microns may appear due to deformation of the plastic film because of wear or the thermal expansion. Compared to glass or quartz masks, polymer films have a relatively short life—they cannot be cleaned effectively, and they tend to bend irreversibly during handling. For designing microfluidic system prototypes, however, lower resolutions and decreased lifetime are not a problem, and the decrease in mask cost can lead to an opportunity to experiment with plenty of designs for the cost it would take to purchase a single quartz mask. 4.3.1.1

Computer-Aided Design (CAD) Tools

Mask data can be transmitted to the mask manufacturer in a large number formats, starting from raster or bitmap drawings, schematics made in illustrator drawing packages, and ending with vector file format such as DXF and GDS II (the standard in the industry). These record all dimensions with infinite accuracy, and allow specification of several layers (the different masks within the design), each with its own polarity, which can usually be interpreted directly by the printers used for mask generation. Many software tools exist for mask design, starting from general engineering design packages, tools

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for electronic circuit and VLSI design, and ending with highly specialized MEMS and microfluidic design toolkits. In the following examples we will use Coventor Ware [19], which provides a complete design, simulation, and vizualisation suite focused on microfluidics and general MEMS. This choice of software package was made for purely illustrative purposes, and many other competing products exist. Mask design in Coventor starts with process definition. The main lithography steps are listed and associated with manufacturing actions such as material deposition or etching. Different etch parameters such as overetching and angle of the side walls can be defined. Each lithographic step introduced in the process definition is associated with a specific CAD layer in the design. The process description introduced in Coventor is relatively simplistic; however, it is sufficient most of the time for creating and visualizing threedimensional structures once the two-dimensional mask design is completed, and therefore provides extremely useful feedback: by watching the 3-D model of the device, the user can rapidly identify errors in the process definition or in the layout itself. Any CAD software provides a number of design primitives that can be used as such or combined to create more complex patterns. Typical primitives involve rectangles, squares, lines, circles, arcs, and ellipses. The dimensions and positioning of the different parts can be specified from the graphical user interface, from the command line, or from a program. Most CAD packages allow Boolean operations to be performed between different primitives on the same layer, or between different layers. This can be very useful for generating relatively complex shapes such as the intersection of two circles. A very powerful feature that is typically provided by CAD programs is a programming interface, which can be extremely useful, particularly when complex geometries that can be described by an algorithm need to be generated. Writing a program allows the user to easily change the parameters of the geometry and see the results rapidly appear in the CAD model as the program is executed. By comparison, performing all the changes manually would take a very long time and would be prone to errors. Programs can be provided with a graphical front end, so that users can invoke a program without being necessarily expert programmers themselves; a graphical window allows them to introduce the device parameters, which are then transfered to the program that executes the commands and builds the geometry. Programming languages used by different CAD packages include C, Tcl-Tk, and Lisp. Another feature of MEMS and microelectronics packages is their use of the cell concept. Cells are components made from different design primitives and bound together in a single object, which can in turn be instanced in other

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Figure 4.9 A typical fluidic design, representing a 4-inch wafer with several instances of the emulsion generator discussed in text. The design used multiple instances of the same cell, and includes dicing marks and alignment crosses.

designs or cells. If a detail within the cell is changed, then all cell instances are automatically updated. Cells can also be instanced as arrays, where the horizontal and vertical spacing of the cells can be specified. This is extremely useful, for example, when several identical instances of a fluidic system need to be assembled on a single mask (an example is shown in Figure 4.9). Individual layers in a design that are associated with lithography steps are called active layers. These are the only layers that are exported into the production file and printed onto masks. The other layers that exist in a design play an indirect role: either they are used as dummy layers in Boolean operations, to obtain certain complex geometries, or they are used as design guides (a substrate layer that includes the outer edges of the mask and the contour of the wafer is typical). Layers are normally rendered in different colors or fill patterns. An example of mask design corresponding to the emulsion generator geometry described earlier in the text shown in Figure 4.7 (left panel) is represented in Figure 4.10. As one can see, the particular rounded shape of all the bends and connections in the fluidic channels is created by a superposition of individual design primitives: rectangles, triangles, circles, and arcs.

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Figure 4.10 An example of mask design for the emulsion generator discussed in the text. Top left: The CHANNEL layer, used for DRIE etching of the fluidic channels, which must be 100 µm deep (except where covered by the TERRACE layer, where the first 97.5 µm are protected by a photoresist layer). Top right: The HOLES and TERRACE layers. Bottom left: superposition of all mask layers. Inset: Detail of the terrace and emulsion-generation channels.

4.3.1.2 Alignment and Dicing Marks

A critical component of any multilayer photolithographic process are the alignment marks. These are features designed on the mask that help in the registration of new masks with the previous lithographic steps. The first lithographic layer typically includes a number of alignment marks equal to the number of subsequent lithographic steps. Each subsequent mask includes a similar alignment mark that needs to overlap with the corresponding mark on the substrate. In the design it must be clearly shown which crosses correspond to which lithographic steps, by explicitly marking the alignment crosses on the first lithography layer with the names of the corresponding subsequent layers. Any ambiguity will lead to delays in manufacturing in the best case (as the person performing the alignment will have to ask for guidance from the designer), and to severe registration errors and shifts with catastrophic consequences in the worst case. The alignment marks must serve two functions. First, they must allow a very rapid rough alignment of the mask with the substrate; to this end, the mask must include patterns that are easy to locate in the mask aligner. Second,

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the alignment marks must contain patterns fine enough to allow very accurate registration to be performed. Typically, the alignment pattern includes several scales; an example is shown in Figure 4.9 where three different sizes of crosses are used—the largest crosses are responsible for the rough alignment, and the smallest ones are responsible for the fine registration. Many different types of alignment marks are used. Normally, the choice of alignment marks used is left to the designer, with one notable exception: robotic mask aligners that use pattern recognition software to perform the alignment automatically require certain patterns that have been preprogrammed in their memory. The cross pattern described above works well for accuracies on the order of the micron, which is typically enough for microfluidics, especially if polymer (transparency) masks are being used. One important remark: The alignment marks used on the first layer need to be different from those used on subsequent layers. In the case of crosses, this implies that the size of the crosses used on the first layer needs to be larger than the crosses on the subsequent layer (for a positive mask) or smaller (for a negative mask). This is required so that the crosses on the substrate are not covered completely by the crosses on the mask (which would lead to the impossibility of further alignment). Figure 4.11 describes this requirement. The alignment marks are normally located on the horizontal diameter of the wafer, close to the edges. For a 5-inch mask (which is the standard size used with 100 mm diameter wafers), the alignment crosses are typically separated by 80 mm. 4.3.2 Fluidic System Optimization Fluidic systems can be surprisingly complex when compared to other types of physical systems. Their behavior may depend on a huge number of variables, and the relevant parameters are usually application-specific. The operation pressure, temperature, fluid type, and flow rate, the physical fluid properties (density and viscosity), the chemical properties, and pressures generated by various actuators and pumps, in addition to many other parameters, may play a critical role in certain applications. It is therefore important to be able to single out the most important phenomena and to model their effects. Surprisingly, simple physical models sometimes provide very useful guidance to the designer, and may offer insight into the observed phenomenology. Methods and software for modeling electronic circuit behavior can also be applied to fluidic circuits. As we have seen in Chapter 3, a close analogy can be made between simple microfluidic channel networks and RC electrical circuits. Fluidic time constants and flow rates can thus be determined by modeling the system as an electrical circuit. Computational fluid dynamics

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Figure 4.11 Alignment crosses on the first and second photolithography layers, in the case of a positive second mask (left) and of a negative second mask (right). The edges of the crosses on the first mask layer must always be visible to allow accurate alignment.

software packages are also available for simulating microfluidic behavior in complex geometries, where analytical approaches fail to provide the required answer. By dividing the fluid volume into finite elements and solving the discretized Navier-Stokes equation on these elements, one can obtain a relatively accurate description of the fluid behavior. The finite element method, however, becomes prohibitively time-consuming for large systems. Examples of these techniques will be given in the following sections, as applied to different microfluidic geometries of interest. These are not intended as a manual for microfluidic simulation techniques, but rather as a complement to the theoretical approaches provided in Chapter 3, and as a guide into the complexities of real-life microfluidics design. 4.3.2.1 Analytical Modeling and Electrical Circuit Equivalence

Let us consider the peristaltic micropump manufactured by multilayer soft lithography, which we described earlier in the chapter. We consider three consecutive valves, which are actuated in the peristaltic sequence (101, 100, 110, 010, 011, 001) that pumps fluid from left to right. We will consider an actuation pressure of Pa = 50 kPa, a valve volume of Vv = 100 × 100 × 10 µm3 =100 pl, a valve closing pressure Pc = 40 kPa and correspondingly a membrane restoring force F = 0.4 mN (corresponding to a pressure of 40 kPa distributed uniformly over the valve area). We will assume the fluidic channel geometry to be 100 µm×10 µm×1 cm for both the inlet and outlet channels,

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Figure 4.12 A schematic of the micropump, showing the three valves working in the peristaltic mode, the inlet and outlet channels (lengths Lin , Lout ), and the valve volume Vv of the middle valve (always shown in the closed position). The top image represents the push phase, and the bottom image represents the pull phase.

and we will consider water as the pumped liquid. These parameters are essentially identical to those published by Quake et al. [16], except for the inlet and outlet channel lengths Lin = Lout = 0.5 cm (which were not published, but which are essential, as we will argue, to determine the maximum pumping rate). We will also consider the inlet and outlet pressures to be equal (pump is functioning under no load). We notice first that there are two distinct pumping phases: a “push” phase, where an upstream valve is fully closed and a downstream valve is being actuated, the inflating valve membrane thus pushing the corresponding volume of fluid towards the outlet, and a “pull” phase, where the downstream valve is fully closed and the upstream valve is released, the relaxing membrane thus pulling the corresponding volume from the inlet (Figure 4.12). In the push phase, the actuation pressure is applied directly to the membrane, and if enough pressure is applied, the fluid can be rapidly ejected from under the membrane (time constant τpush ). To model this, the membrane must be considered a nonlinear element, which is fully compliant and provides an elastic pressure proportional to the total displaced volume dp = dV /χ for V < Vv (corresponding to χ = Vv /Pc ), but becomes completely rigid for V ≥ Vv (χ → 0). By contrast, in the pull phase the membrane relaxes under its own elastic force, the dynamics being exponential and much slower (time constant τpull ). Each push-pull sequence results in a volume Vv being displaced from the left of the pump to right. The maximum flow rate achievable with such a pump is therefore roughly given by the valve volume divided by the sum of the inflation and deflation time constants (as that is the minimum time one needs to wait for a complete push-pull sequence to be achieved): Φ = Vv /(τpush + τpull ). To calculate the push time τpush, we assume that Pa > Pc . The

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membrane will inflate as an elastic medium until V = Vv (the membrane is fully inflated, corresponding to the closed valve position), and then it will not change its volume anymore. The time τpush it takes for the membrane to inflate can be calculated by an electrical analogy with the RC circuit represented in Figure 4.13 (a capacitor charging through a resistor). For this we will remember the analogies made in Chapter 3—the compliance of any fluidic component (χ) is equivalent to a capacitance C, and the hydrodynamic resistance of the outlet channel (ρout ) corresponds to a resistance R. The valve volume Vv corresponds to a maximum charge Qmax allowed on the capacitor: when the charge on the capacitor reaches this value, it does not increase anymore. The actuation pressure Pa corresponds to the voltage applied by the voltage source V , and the valve closing pressure Pc corresponds to the voltage on the capacitor when Q = Qmax . It is trivial to calculate the charge on the capacitor as a function of time: t Q = CV (1 − exp(− RC )), which yields τpush as the solution to the equation Q = Qmax : Qmax ) (4.1) τpush = −RC ln(1 − CV By using the analogy between the fluidic and the electrical circuit outlined in the text, we obtain: τpush = −ρout χ ln(1 −

Vv ) χPa

(4.2)

which, taking into account the value of χ = Vv /Pc , becomes: τpush = −

Pc ρout Vv ln(1 − ) Pc Pa

(4.3)

In the limit of strong actuation (Pa ≫ Pc ), we perform a Taylor Vv expansion to obtain, to first order: τpush ≈ ρout Pa —the actuation time during the push phase is inversely proportional to the actuation pressure (i.e., it can be reduced significantly by driving the system with a high pressure). It is now a simple exercise to calculate τpull , the relaxation time during the pull phase. Electrically, it is equivalent to a capacitor discharging through a resistor. This results in an exponential decay, with the time constant being given by: ρin Vv τpull = ρin χ = (4.4) Pc We can use the formula we learned in Chapter 3 to calculate the 12ηL hydrodynamic resistance of the channels: ρ ≈ h3 w(1−0.63 h , with L = Lin = ) w

Lout = 0.5 cm, h = 10 µm, w = 100 µm and η = 1 mPa s. With these

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Figure 4.13 (a) The micropump “push” phase is equivalent to the charging of a capacitor in an RC circuit. The corresponding solution (b) can be used to estimate the time τpush when the maximum charge Qmax is reached (corresponding to the maximum membrane deflection).

parameters, we obtain τpull = 160 ms and τpush = 128 ms, which leads to a maximum pressure of approximately 3.5 Hz for linear pump behavior, and to a maximum flow rate in the linear regime of 0.35 nl/s. The results published by Quake et al. [16] indicate that the pump operates in the linear regime up to a frequency of about 60 Hz, and delivers 1.5 nl/s at that frequency. While this is a relatively large discrepancy, we can see that the simple calculations we made allow us to provide a rough order-of-magnitude estimate of the peak pump operation parameters. It is possible that the cause of the discrepancy lies in the geometry used by Quake et al., which was not fully disclosed. In particular, the length and geometry of the inlet and outlet channels leading to the device were not published. If we assume that outside the pumping region the height of the channel is 25 µm instead of 10 µm, we obtain a much better estimate for the maximum frequency in the linear regime (48 Hz) and for the flow rate obtained in these conditions (4.9 nl/s). 4.3.2.2

Computational Fluid Dynamics (CFD) Software

Complicated fluidic geometries are difficult to model analytically or by analogy with electrical circuits. In this case, one must rely on numerical calculations—the model needs to be broken in finite elements, the NavierStokes equation needs to be discretized, and the problem is solved by successive (converging) iterations. The procedure for setting up a finite element analysis is relatively straightforward on paper. It starts with defining and meshing the geometry. Many problems can be meshed using twodimensional meshes; this is true, in particular, of situations where the system

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possesses a certain type of symmetry (rotational or translational) so that the equations do not depend on the corresponding coordinate. This is the ideal case, since two-dimensional problems are much easier to solve numerically. The most often used two-dimensional meshes are triangular and quadrilateral. Whenever the problem at hand cannot be reduced to two dimensions, a full three-dimensional simulation needs to be performed. The 3-D simulations are much more computationally intensive than their 2-D counterparts, and therefore they require a much longer simulation time. Even if the problem cannot be reduced to two dimensions, often it can be simplified by observing certain symmetries of the geometry—a mirror symmetry about a plane can reduce computation times by half. Two mirror symmetries will reduce it by three quarters. Regardless of problem dimensionality, and regardless of the software used for simulating the problem, the mesh must be generated correctly for the simulation code to work well. The quality of the mesh is the single most important parameter when solving a fluidic problem numerically. The mesh should be refined in the regions where high gradients in velocity (or another parameter of interest) are expected, such as around sharp corners. The mesh size should be decreased progressively and the results of the numerical simulation examined. If the results vary significantly as the mesh is reduced, most likely the simulation is not yet accurate enough and the mesh needs to be reduced further. On the other hand, once the results seem to converge and to fall within the desired accuracy window, there is little use in further reducing the mesh size. After a model is properly meshed, one must specify the boundary conditions. Typically, all walls are defined as zero velocity boundaries, in order to verify the no-slip boundary condition. On the other hand, if there is contact between several fluids, then the boundary conditions need to be defined to balance normal and shear stresses at the boundary. The next step in setting up the simulation is to define the initial conditions (the initial fluid velocities, pressures, and so forth). Finally, one needs to decide whether the transients are important (i.e., whether time-dependent behavior before the system reaches a steady state needs to be simulated), or whether a a steadystate solution is sufficient. As an example of computational fluid dynamics simulations, we will look at pressure-driven flow in a shallow channel having sinusoidally modulated lateral walls. We will use Coventor again, as it allows the microfluidic geometries to be imported from the CAD software straight into the CFD simulation tool. We will use quadrilateral extruded bricks as the mesh element as shown in Figure 4.14 (top left), and we will apply no-slip boundary conditions on all the walls except the inlet and the outlet, where the pressures

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Figure 4.14 CFD simulation results. Top left: Pressure field overlaid on the 3-D mesh. Top right: Slices showing contour plots of longitudinal velocity, at several longitudinal positions. Bottom left: Stream lines in the symmetry plane of the device. Bottom right: The 3-D velocity vector field at the exit of the device. (Courtesy of M. Stoffel, ESIEE.)

will be specified. The simulation provided the steady-state results shown in the different panels of Figure 4.14. As it is apparent from this section, CFD simulations can provide very detailed results about the behavior of a fluidic system. From this point of view, they are a critical tool for the microfluidic designer. CFD simulations, however, always need to be approached with care—a small error in the simulation setup, or lack of convergence in the iterative solving process, can lead to completely erroneous results. It is therefore a very good idea to validate the results of the CFD simulations, whenever possible, against analytical results.

4.4 Conclusions This chapter has reviewed the main aspects of microfluidic design: functional diagram, choice of integration options, choice of technology, definition of the

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fabrication process, mask design, and system simulation and optimization. Examples are given to illustrate the important points. After reading this chapter, the reader should have an idea of what a complete design project implies, and how to start thinking about it. The focus is intentionally on breadth rather than on depth: this chapter is not intended as a specialized manual in any one technique, but rather as an overview of techniques commonly used. This chapter provides the basics; references are given to more detailed approaches about each specific topic. The reader is expected to experiment hands-on with design and simulation tools, as well as with different fabrication processes and technologies. It is important to try, to fail, and to retry. Success is reserved to those who persevere.

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References [1] Plummer, J., Deal, M., and Griffin, P., Silicon VLSI Technology: Fundamentals, Practice and Modeling, Upper Saddle River, NJ: Prentice Hall, 2000. [2] Nguyen, N., and Wereley, S., Fundamentals and Applications of Microfluidics, Norwood MA: Artech House, 2002. [3] Hsu, T., MEMS & Microsystems: Design and Manufacture, Boston, MA: McGraw-Hill, 2002. [4] Maluf, N., and Williams, K., Introduction to Microelectromechanical Systems Engineering, Norwood MA: Artech House, 2004. [5] Groover, M., Fundamentals of Modern Manufacturing: Materials Processes, and Systems, New York, NY: John Wiley and Sons, 2007. [6] Beeby, S., Ensell, G., Kraft, M., and White, N., MEMS Mechanical Sensors, Norwood MA: Artech House, 2004. [7] Ehrfeld, W., Hessel, V., and Lowe, H., Microreactors: New Technology for Modern Chemistry, Weinheim, Germany: Wiley-VCH, 2000. [8] Mack, C., Fundamental Principles of Optical Lithography: The Science of Microfabrication, West Sussex, England: Wiley-Interscience, 2007. [9] Suzuki, K., and Smith, B., Microlithography: Science and Technology, Boca Raton FL: CRC Press, Taylor & Francis group, 2007. [10] Chakrabarty, K., and Zeng, J., Design Automation Methods and Tools for MicrofluidicsBased Biochips, Dordrecht, The Netherlands: Springer Verlag, 2006. [11] Griebel, M., Dornseifer, T., and Neunhoeffer, T., Numerical Simulation in Fluid Dynamics: A Practical Introduction, Philadelphia, PA,: Society for Industrial and Applied Mathematics, 1998. [12] Tu, J., Yeoh, G., and Liu, C., Computational Fluid Dynamics: A Practical Approach, Oxford, UK: Elsevier, 2008. [13] Stolarski, T., Nakasone, Y., and Yoshimoto, S., Engineering Analysis with Ansys Software, Oxford, UK: Elsevier, 2006. [14] “Cornell Nanoscale Science and Technology Facility Brochure.” http://www.cnf. cornell.edu/doc/cnfmems.pdf. [15] Quake, S., and Scherer, A., “From Micro to Nanofabrication with Soft Materials,” Science, Vol. 290, 2000, p. 1536. [16] Unger, M., Chou, H., Thorsen, T., Scherer, A., and Quake, S., “Monolithic Microfabricated Valves and Pumps by Multilayer Soft Lithography,” Science, Vol. 288, 2000, p. 113. [17] Anderson, J., Chiu, D., Jackman, R., Cherniavskaya, O., McDonald, J., Wu, H., Whitesides, S., and Whitesides, G., “Fabrication of Topologically Complex ThreeDimensional Microfluidic Systems in PDMS by Rapid Prototyping,” Analytical Chemistry, Vol. 72, 2000, p. 3158.

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[18] Sugiura, S., Nakajima, M., Iwamoto, S., and Seki, M., “Interfacial Tension Driven Monodispersed Droplet Formation from Microfabricated Channel Array,” Langmuir, Vol. 17, 2001, p. 5562. [19] Coventor, Inc.: Cary, NC, http://www.coventor.com.

5 Integrated Microfluidic Systems

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5.1 Introduction As it has become apparent from the previous chapters of this book, there is a natural tendency towards miniaturization and integration in microfluidic devices. There are multiple underlying reasons: the economic aspect of manufacturing is the main driver in most cases, as small integrated devices made in batch processes require fewer external components, less time to manufacture, smaller quantities of raw materials, and less manual assembly work (all of which imply costs). Size is often a driver as well—smaller devices can enable deployment in new environments, or can allow multiple measurements to be performed in parallel, within less volume than a traditional measurement. There are additional benefits to the end user—by using integrated devices that use less sample volume and less reagents, it is possible to reduce either the cost or the time per analysis (which can both be very important economic criteria in many applications). This chapter attempts to explore a diverse array of examples of integrated microfluidic devices, which are close to being or already are being commercialized. The examples assembled in this chapter illustrate different approaches to microfluidic integration, by selecting examples of similar technologies applied in completely different application areas, or, on the contrary, examples of applications competing in scope but using very different technologies. We will examine application areas as diverse as biotechnology, drug delivery, gas chromatography, and fluid measurements and analysis. The technologies covered will include both continuous fluid flow and digital (droplet-based) aproaches, whereas the manufacturing methods illustrated will cover fabrication in soft elastomer, thermoplastics, and silicon/glass materials. It is important to note that this chapter does not make a comprehensive analysis of the industry, but rather describes a few specific examples that were chosen based on their relevance to the theme of this book, and on their capability to reflect the diversity of applications and technologies that are being applied in this area. Many of the applications presented are in the emerging phase, whereas others have been existing on the market for a number of years. The mix is quite heterogenous, and purposedly so, as it is a direct reflection of the current state of the microfluidics market. The purpose of this chapter is merely to provide an initial exposure of the reader to the complexity and diversity of the microfluidics market and to provide a number of commercial examples where integration of multiple microfluidic functionalities, or of microfluidic technology with other technologies such as CMOS or MEMS, has been achieved succesfully. The choice of specific examples in this chapter does not represent in any way an endorsement by the


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author, nor does it represent an advice to select one technology versus another in any specific application.

5.2 Remarks About the Microfluidics Market As early as 2004, microfluidics was named in the press as one of the 10 emerging technologies that will change our world [1]. Such statements are being made increasingly often today, as the technology is starting to deliver to its promise. We are starting to see research moving from academic and government laboratories to the marketplace. Start-up companies are sometimes initiated by academic researchers, and more often by their graduate students, who are excited to continue developing their thesis research work with the motivation of a commercial market-relevant application in mind. Currently, universities and public institutes in several countries strongly support such entrepreneurial actions: startups are being created around technologies that are typically licensed from the universities in the initial phase, then further developed for commercial applications. Academic institutions often facilitate access to competent legal counsel and to funding avenues such as venture capital, for example, by organizing periodic technology meetings that bring together members of the academic community and business angel investors. The interest is on both sides, as many investors realize the fast growth potential of emerging technologies and are ready to support some of them in the start-up phase. Country and local governments are equally keen in supporting innovative technology projects, many grant schemes being initiated to facilitate the transfer of innovative research in microsystems and microfluidics to the marketplace. Large corporations have also realized the benefic potential of these technologies to their businesses; often, joint laboratories are being created between academic institutions and industrial partners. These are mutually advantageous, as corporations have a chance to guide the research toward their market goals, while academic research groups receive some funding for their research and often discover unexpected practical relevance of their work. Governments generally offer financial support for such joint laboratories, often very generously. Some corporations are developing innovative tehnologies in-house, and integrate them directly in their line of products; however, microfluidic devices have been less often in this position than standard MEMS products. In such an active and fertile environment, many technologies attempt the transition from the research lab to the marketplace— and some of them become quite successful. Several metrics of the activity in this area can be

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Figure 5.1 The number of publications and granted U.S. patents on the topic of microfluidics, as a function of publication year in the last decade. The 2010 numbers were extrapolated from data available on July 1, 2010 (author’s research, based on the Web of Science and the USPTO databases).

considered: the number of publications involving the topic of microfluidics, for example, evolved continuously from 118 in 2000 to 980 in 2009, with a predicted number of 1,072 publications in 2010 (extrapolation based on July 1, 2010, data from the Web of Science database; Figure 5.1, left panel). The number of granted U.S. patents involving microfluidics during the same period (Figure 5.1, right panel) shows less dramatic development. This discrepancy between the two graphs is understandable, as patents are awarded only several years after publication due to the long examination process. Also, each granted patent may lead to several individual publications. It is to be expected that, as the number of commercial applications increases, the patent numbers will explode in the next few years, as suggested by extrapolated 2010 data. Another interesting phenomenon is the partition of patents granted in microfluidics between academic laboratories, businesses (both large and small, established or start-ups) and individuals. Academic laboratories accounted for 19% of patents between 1999 and 2006, businesses for 80%, and individual inventors for a mere 1% [2]. This situation is likely to evolve, as academic institutions realize the enormous advantage of possessing an IP portfolio, and therefore are becoming a lot more aggressive in patenting new technologies developed in-house. The geographical distribution of granted U.S. patents in the area of microfluidics is heavily biased in favor of U.S. businesses and laboratories. As many as 95% of U.S. patents related to microfluidics between 1999 and 2006 were granted to U.S. institutions [2]. This is normal to some extent— U.S. institutions are very reactive to technologies with strong market potential, and are backed by a strong venture capital community that shortens the lab-tomarket delay, but requires a strong IP position as a prerequisite. Companies


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from other countries may tend to patent locally first, and then extend the patent worldwide (including in the United States), which is a relatively long process. The most active countries in developing integrated microfluidicsbased products are the United States on the American continent, Switzerland, Germany, and the Netherlands in Europe, and Japan and Singapore in Asia. It is also interesting to note that a lot more publications tend to come out of research performed in microfluidics at U.S.-based companies, as compared to European companies, for example. While the underlying reasons for this situation can be the subject of many debates, the facts show that U.S.-based companies tend to have stronger IP positions and, therefore, to feel more comfortable openly discussing their technologies and results in the scientific literature. This discrepancy is probably going to change in the next few years, as European governments and academic institutions are introducing strong incentives for companies to develop innovative technologies, often in collaboration with academic laboratories.

5.3 Selected Commercial Technology Examples In this section we will review a number of companies that developed technologies capable to integrate multiple functionalities on a single microfluidic chip or on a hybrid device. For each company we will provide a brief history, describe its main products, target markets, and patent portfolio, and review the technology involved and the approaches taken towards integration. A brief summary of the different companies analyzed in this short review is given in Table 5.1. 5.3.1 Microfludic Large-Scale Integration: Fluidigm 5.3.1.1 History

Fluidigm Corporation (originally called Mycometrix) was founded in 1999 to commercialize microfluidic large-scale integration (mLSI) technology capable of creating complex multilevel soft lithography devices (branded under the name Integrated Fluidic Circuits (IFC)). The technology was born in the laboratories of Steve Quake, a professor at the California Institute of Technology at the time, Fluidigm cofounder, and scientific advisory board member. The first commercial product using IFCs was the Topaz system for protein crystallization in 2003, followed by BioMark in 2006—a DNA amplification system allowing extremely high throughput, and FLUIDIGM EP1—a system that allows rapid and economical genotyping, unveiled in 2008. In 2005 Fluidigm opened a large microfabrication facility in Singapore, specializing in IFC fabrication.

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Table 5.1 Selection of Companies Integrating Microfluidic Technology in Their Products.

Company Name Fluidigm Corporation San Francisco, CA (United States)

RainDance Technologies Lexington, MA (United States)

ThinXXS Microtechnology AG, Zweibrücken (Germany) ISSYS (Integrated Sensing Systems), Ypsilanti, MI (United States) Debiotech SA, Lausanne (Switzerland)

Sensirion AG, Staefa (Switzerland) C2V, Concept to volume, Enschede (the Netherlands)

5.3.1.2

Selected Companies Approach to Integration Multilevel PDMS devices with integrated fluid manipulation capabilities (hydraulic valves, pumps); integrated fluidic circuits PDMS devices capable to manipulate (generation, splitting, coalescence, mixing, sorting) and analyze microreactor water droplets by fluorescence Plastic devices manufactured using precision molding and incorporating piezoelectric pumps, valves, and custom channel geometries MEMS silicon/glass devices incorporating silicon microtubes and high-vacuum packaging MEMS silicon/glass high performance piezoelectric micropumps with integrated pressure sensors; microneedle arrays; biocompatible coatings Combined on-chip CMOS and sensing unit Silicon/glass and polymer manufacturing; Multicomponent hybrid integration (fluidic multi chip module); microvalve and micro TCD technology

Market Applications Protein crystallizatiom; genotyping; DNA analysis; PCR

Genomics; DNA analysis; PCR

Custom manufacturing; component kits; development of hybrid plastic/silicon integration platform; joint project development Implantable sensors; drug delivery; fluid analysis; chemical analysis ; pressure sensing; foundry services Drug delivery pumps; implantable medical devices;

Gas/liquid flow sensing; differential pressure; humidity sensors Micro gas chromatography; custom manufacturing; hybrid integration of building bricks

Technology

Fluidigm uses a technology branded MSL, and previously referred to as mLSI [3], to fabricate IFCs, by analogy with the corresponding term from electronic integrated circuits, very large-scale integration (VLSI). The technology was


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Figure 5.2 The microfluidic valve density using mLSI technology was initially growing four times faster than Moore’s law predictions for integrated curcuits. While it is not clear how long such exponential growth can be maintained, the data shown in these graphs is impressive and encouraging. (Inspired from [4].)

described in Chapter 2, to which the reader is referred for further technical details and references. mLSI involves multiple-layer soft lithography devices, with embedded flexible membranes and hydraulic control channels, enabling complex microfludic infrastructures integrating thousands of valves and pumps to be manufactured on a single chip. The degree of integration that can be achieved with this technology is impressive, and it appears that the total number of valves and pumps that can be built per unit surface grows exponentially, with an initial time constant that is significantly faster than that predicted by Moore’s law for integrated circuits (Figure 5.2). Using mLSI technology, both valves and peristaltic pumps can be built, which allows a large number of individual operations to be perfored on nano-liter scale fluid volumes. A cell sorter unit using multilayer soft lithography has been implemented, capable to detecting and separating certain types of cells based on the presence or absence of a fluorescent signal [5]. Similarly, a cell culture systems integrating functional blocks for injecting the cells, pumping fluids through the system, performing mixing operations, and multiplexing the fluidic path to multiple culture chambers is shown in Figure 5.3 [6]. Very large numbers of such specialized units that work in parallel

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Figure 5.3 An example of a cell culture chip realized in mLSI technology. Colored water was used to mark different portions of the device, which contains culture chambers, multiplexers, valves, peristaltic pumps and a mixer. (Reproduced with permission from [6]. Copyright 2007 American Chemical Society).

can be assembled and controlled using a relatively small number of hydraulic control lines. One important aspect in microfluidics involves multiplexing and demultiplexing fluid paths—selecting which one of several fluidic inputs is connected to an output or vice versa. This can be used for any number of applications, such as for cell sorting based on the presence of certain characteristics, or for selecting which of several chemical reagents is combined with a given sample. mLSI technology solves this problem in an elegant way, by using hydraulic actuation pressures that allow the flexible membrane belonging to a wide actuation channel to completely block the corresponding fluid channel, but will not allow channel blocking for narrow actuation channels. By using a binary scheme with multiple parallel actuation channels that can block half of the channels each, it is possible to control 2n fluid channels using 2n control lines (an example for n = 3 is shown in Figure 5.4). The technology has a number of great advantages for biotech and pharmaceutical applications: by using soft materials like silicone elastomers (PDMS) to fabricate the tiny flexible valve actuation membranes, it enables


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tens of thousands of fluidic manipulations and chemical or biological reactions to be performed in a parallel manner, thus cutting cost, saving time, and reducing reagent usage. PDMS is biocompatible and optically clear, thus allowing real-time visualisation of the fluidic processes. mLSI is very versatile, and so far it is ahead of any competing technology in single-chip integration density of fluidic components. Some shortcomings of the technology stem from the same roots as its advantages: by using a soft material for manufacturing, the devices cannot operate at very high pressures, which is a requirement for certain industries. Silicones like PDMS tend to be peremeable to gases and to swell in contact with a number of organic solvents, which severely limits their usage in certain applications. The chemical compatibility issue may be eliminated by the development of other soft materials with lower permeability and higher solvent resistance. Fluoropolymers have been investigated to this end [7]. The pressure capability of the devices may be improved by using stronger materials for fabricating the fluid and actuation channels (e.g., by bonding or clamping two rigid channel plates with a thin membrane sandwiched in the middle). Other shortcomings are not as easy to eliminate, since they stem from the physics of device operation. By using pressure-driven flow in the microchannels, mLSI technology is prone to dispersion issues due to the parabolic nature of the flow (e.g., Taylor dispersion). This tends, on one hand, to limit the resolution and the time response of the device in certain analytical applications and, on the other hand, to increase the flush time required to clean the device. It is not clear if the technology can be used in conjunction with other types of flow (e.g., of electrokinetic nature) to reduce the dispersion issue.

5.3.2 RainStorm Droplet Technology: RainDance 5.3.2.1 History

RainDance Technologies was founded in 2004 by a number of scientists from the United States, France, and the United Kingdom, to commercialize dropletbased microfluidic technology originally developed primarily in Dave Weitz’s laboratory at Harvard University. The company has an impressive scientific board including three Nobel laureates, and it has managed to grow relatively fast thank to a combination of government grants and private financing. Its flagship product is the RTD1000, which uses the microdroplet technology branded RainStorm to perform DNA sequencing and accelerate the targeted sequencing of the human genome.

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Figure 5.4 Left panel: Operation of a 1:8 multiplexer, using 6 control lines. To select channel 3, one first needs to block channels 4–7 (by actuating the second channel from the top corresponding to bit3=0). This leaves channels 0–3 open, so next one needs to block channels 0–1 (by actuating the third channel from the top corresponding to bit2=1), which leaves open channels 3 and 4. Channel 4 is blocked by actuating the fifth channel from the top, corresponding to bit1=1. Right panel: Example of line and column multiplexers used to address one of 1,000 fluidic (25 × 40) chambers. (Reproduced from [3].)

5.3.2.2

Technology

RainDance uses a droplet technology branded RainStorm to perform a number of basic fluid manipulation operations in sequence on the same chip. Droplets can be created, charged, neutralized, coalesced, mixed, split, sorted, and optically interrogated. Each droplet is essentially an independent chemical reactor surrounded by an inert fluorinated oil, which assures that the droplets are never in contact with air or with any contaminants coming from the channel walls. Droplet creation is achieved using microfluidic flow focusing [8], a process by which a stream of water is surrounded on both sides by streams of oil, leading to a periodic capillary instability and thus to the formation of monodisperse droplets. By using an electric field in addition to flow focusing, the size of the drops can be electrically controlled—the formation of a Taylor cone can drastically reduce the droplet diameter [9, 10]. These processes are shown schematically in Figure 5.5. In addition to this control over the droplet size, the application of an electric field also results in the formation of electrically charged drops, which can then be manipulated by electrostatic forces. In particular, when a voltage is applied between two opposite flowfocusing junctions, the respective droplets acquire opposite charges and their


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Figure 5.5 Left: A schematic flow focusing geometry, whereas the water stream coming from the left is split into droplets by the two lateral streams of oils. The stability of the droplet production and the size and uniformity of the droplets depends on a number of parameters, such as device geometry, flow rates involved, the surface tension of the fluids and types and concentrations of the surfactants used. Right: the same schematic geometry, but with an electric field applied, results in charged droplets with voltage-controllable size.

formation becomes synchronized, resulting in a tendency of the droplets to coalesce and neutralize the charges. When the electric field is sufficiently strong, coalescence is forced in 100% of the cases, which creates a convenient technique to join droplets [9]. This can be applied, for example, to combine a droplet containing a sample to other droplets containing different chemical reagents. Drop sorting can be achieved using electrical fields for electrically charged droplets, or using dielectrophoretic forces for manipulating neutral drops. The magnitude of the dielectrophoretic force depends on the dielectric constant contrast between the droplet and the surrounding fluid (which is very high in the case of water drops in fluorinated oil), and on the electric field gradient. By carefully designing electrode geometries, the field gradient can be optimized, resulting in a convenient way to sort droplets electrically (Figure 5.6). Upstream from the sorting junction, the droplets can be interrogated optically and a decision can be made to send the droplet in either one of the two outlets based on the result of the interrogation. In one application, cells are enclosed within droplets and a fluorescent label is used that can distinguish between cells possessing certain DNA characteristics. The cellenclosing droplets can then be sorted based on the presence or absence of the fluorescence signal [11].

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The RainStorm technology has a number of advantages, namely, the physical separation of the different droplets, which allows them to act as completely independent reactors and to avoid dispersion and wall contamination issues. Droplet libraries, containing different chemical or biological content, can be created and stored for later usage. The technology allows very rapid manipulations to be performed, at the rate of several thousand drops per second, which leads to an enormous throughput in a large number of applications. The technology is rather versatile from a manufacturing point of view: as it does not fundamentally depend on the mechanical properties of the materials used, it can be implemented in elastomeric, plastic, or silicon/glass materials. The only requirement is transparency, for optical interrogation. There are a number of shortcomings of the technology as well: it seems that, while being suited very well to sorting applications where one requires YES/NO types of answers (as obtained from the presence or absence of fluorescent signal), it seems that quantitative measurements may be more difficult to obtain from a single droplet. The fluidic design and tolerances during manufacturing need to be well controlled to assure robust operation, as the performance of dielectrophoretic separation depends strongly on hydrodynamic forces, and hence on the droplet concentration in the outlet channels (there is therefore a certain “memory” of the device). Therefore, to assure robust operation, all such sorting operations need to be performed in very dilute conditions, which limits throughput to some extent. Finally, it is not clear whether all droplet operations might be possible with other types of fluids than water, since many of the operations depend on the conductivity and dielectric constant of water. 5.3.3 Modular Plastic Microfluidics: ThinXXS 5.3.3.1

History

ThinXXS was founded in 2001 as a spinoff from the Institute of Microtechnology Mainz (IMM) by Lutz Weber, the current ThinXXS chief technology officer. Its core technology is based on high-precision toolmaking and plastic injection molding, which it uses for the microfabrication of microfluidic devices and micropumps. ThinXXS manufactures and sells one of the few existing commercial micropumps, a microfluidic kit including several modular building blocks, and also offers custom product development. It has made a number of partnerships to codevelop products for the pharmaceutical (drug discovery) and medical (blood sampling) industries. It was involved in the µBUILDER project sponsored by the European Union, where it was


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Figure 5.6 Dielectrophoretic droplet sorting technique. (a): The top view of the microfluidic device; (b): The cross-sectional view; (c): if the device is operated with no electric field, all drops move to the left due to the lower hydrodynamic resistance of that channel; and (d): with an electric field applied, all the drops are guided to the right. (Reproduced with permission from [12]. Copyright 2006, American Institute of Physics.)

responsible for the integration of silicon/glass sensors with economical plastic fluidic components and pumps. 5.3.3.2 Technology

ThinXXS Microtechnology develops and manufactures disposable microfluidic devices from plastics for the diagnostic, pharmaceutical, analytical, and medical industries [13, 14]. ThinXXS also specializes in customerspecific solutions for applications in immunology, clinical chemistry, DNA analysis, and cell-based research. The core technology of ThinXXS is precision molding of transparent plastics such as cyclic olephin copolymer (COC), and on microassembling and bonding plastic or hybrid parts using laser, adhesive, and solvent-mediated processes. Molding is achieved using specialized injection machines and extremely high-precision molding tool inserts realized in hard alloys using high-speed micromilling. This technology results in submicron accuracy on the final chip and channel dimensions. A number of microfluidic devices, with complex geometries and high tolerances, can be manufactured using such technology (Figure 5.7, left). ThinXXS manufactures one of the few available commercial microfluidic pumps using plastic molding technology (to fabricate membrane and static check valve parts) and hybrid integration with piezoelectric disk actuation. Along with the pump, ThinXXS also offers a number of premade microfluidic slides having the dimensions of a typical microscope slide, which incorporate a number of basic fluidic functions such as mixing and

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Figure 5.7 Left: Customized lancet disk for diabetes testing manufactured using highprecision injection molding of plastics. (Reproduced from [15].) Right: The ThinXXS construction kit allows a number of microfluidic slides integrating pumps and other fluidic elements to be connected into a hybrid system. (Reproduced from [16].)

splitting fluid streams. Custom channel geometries can be realized, and fluidic connectors are available to connect any number of slides (Figure 5.7, right). ThinXXS took a central position in the European microBUILDER project [17, 18], which grouped nine European partners with different competences in microfluidics with the aim of creating a unified platform for microfluidic development. The goal of the project was to develop a hybrid integration technology, allowing complex silicon devices to be integrated with economical plastic microchips and pumps to build functional lab-on-achip microsystems. The project aimed to create a manufacturing consortium, allowing clients to access (via consortium service suppliers) technologies as diverse as silicon multiwafer micromachining, plastic injection molding, and chemical functionalization. The microfluidic slide concept developed by ThinXXS was chosed as the development platform, and the feasability of the technology was demonstrated by integrating silicon flow metering chips onto plastic microfluidic slides [19] (Figure 5.8). The plastic injection technology developed by ThinXXS is very appealing due to the relatively low fabrication cost per piece that could be achieved for high volumes. The microfluidic slide concept allows integration and custom development of a number of fluidic units, with the highest value being added by the possibility to integrate in the design the commercial ThinXXS micropumps. The demonstrated possibilities to hybridly integrate microfluidic chips made using other types of technologies (most notably silicon/glass) via a custom plastic channel plate has strong potential, allowing the lab-on-chip developer to concentrate on the design of mission-specific parts using the most appropriate technology while leaving pumping and fluidic connectivity duties to the low-cost plastic components.


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Figure 5.8 A microfabricated silicon-based flow sensor along with conditioning electronics, integrated on the ThinXXS microfluidic slide platform. (Reproduced from [19].)

There are shortcomings of the technology as well. The plastic materials have chemical compatibility issues—COC, for example, the material of choice for ThinXXS components, has very little resistance to aliphatic hydrocarbons. While the modular microfluidic slide concept is appealing, the connections between chips are relatively bulky, with high dead volumes. It is not obvious to integrate electrical connections on the plastic parts in the injection molding process, so that complex integrated systems requiring electrical interconnects between different components may be difficult to achieve. Finally, the economic advantage of plastic molding may be outweighed by the additional microassembly and bonding operations that are required—the cost advantage of plastic molding versus batch micromachining should be carefully weighed for each project. 5.3.4 Resonant Silicon Microtubes: ISSYS 5.3.4.1 History

Integrated Sensing Systems (ISSYS) was founded in 1995 as a spinoff from the University of Michigan in the United States, to transfer and commercialize MEMS development work by its founders, Nader Najafi, Ken Wise, and Khalil Najafi. The core fabrication technologies allowing innovative MEMS development at ISSYS are the dissolved wafer process (DWP), for which it acquired exclusive rights from the University of Michigan, and microtube fabrication technology using plasma etching. The technology allowed ISSYS

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Figure 5.9 MEMS microtube unit, incorporating electrodes for capacitive actuation and detection, and nanogetters for vacuum packaging. (Reproduced from [20] and [21].)

to develop a line of high-accuracy pressure sensing products, as well as microfluidic systems based on monocrystalline resonating silicon microtubes. These systems allow accurate pressure, density, and flow measurements to be performed in a single microfabricated device, and have led to a number of products developed in-house or within strategic partnerships, such as a fuel concentration sensor, drug infusion monitoring devices, industrial density and flow meters, and wireless pressure sensors. ISSYS is at the point where it is starting to generate revenues from its commercial products in addition to those from its foundry services. 5.3.4.2

Technology

Patented MEMS processes allow ISSYS to manufacture suspended single crystal silicon microtubes, which can be embedded in a microfluidic device along with electrostatic actuation and detection. The technology is based on plasma etching: a channel is etched in a first silicon wafer using a plasma etch step, which is then bonded to a second wafer. The two wafers are thinned and the channel outer wall is defined by a second plasma step to generate a fully suspended, U-shaped, silicon tube. The advantage of the plasma process as compared to other technologies for building microtubes such as those based on highly doping the silicon, resides in the improved versatility and control over the wall thickness. In particular, thick-walled tubes can be manufactured that can withstand pressures up to 40 bar [20]. The silicon microtubes are bonded to a glass substrate on either side, allowing complete encapsulation. Electrodes are patterned on the glass substrates to allow electrostatic actuation (driving the tube in resonance) and to sense its motion. To allow a sharp resonance, the device needs to be operated in vacuum, which is a second technological achievement of ISSYS


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(and its subsisdiary Nanogetters); by encorporating a reactive getter within the resonant vacuum chamber for the tube, it was able to bring the pressure down to below 1 mTorr, which resulted in a resonant quality factor Q for the microtube of order 10,000 to 60,000 [20–22]. One obvious aplication of a high Q-factor fluid-filled resonator is in measuring density—the mass change leads to a shift in resonant frequency that can be detected by appropriate electronics. Very accurate density measurements can be achieved (down to 10−4 relative error [20]), with applications in several areas, such as for accurate detection of fuel concentrations (methanol in water, for example), or for detecting dosage errors during IV drug infusion. By detecting the out-of-plane vibration of the silicon microtube using multiple capacitive detectors, ISSYS has been capable to use the same technology to perform Coriolis-based flow measurements with an accuracy of 0.5%. Temperature can be recorded using the metallic traces on the chip as RTDs, and finally, by detecting the Q-factor degradation (or the damping) due to the viscosity of the fluid within the microtube, it has been possible to use the same device to measure the viscosity of the fluid with errors in the percent range [23]. The resonant microtube technology reveals itself as being very powerful, allowing multiple measurements to be performed with a single device. While the efforts at ISSYS have not necessarily been at integrating multiple types of devices on a single chip, they manage to integrate multiple measurements with impressive accuracy into a single miniaturized device. Being fabricated entirely in silicon, the device is chemically inert for the majority of applications. It can withstand high pressures (albeit with an accuracy loss due to the thicker walls involved). A drawback of the technology resides in the fact that its very specific and relatively nonstandard fabrication procedure does not allow it to easily integrate in a monolythic fashion with other types of technologies. Optical measurements, in particular, are difficult due to the opaque nature of the silicon material. Accurate pressure measurements, another specialty of ISSYS, are also difficult to integrate with the microtube technology. However, the unit could become part of a hybrid platform, integrating many other types of measurements. 5.3.5 Nanopump MEMS Drug Infusion: Debiotech 5.3.5.1 History

Debiotech SA was founded in 1990 and is based in Lausanne, Switzerland. It is closely associated with the the Center for Microfabrication at EPFL (Ecole Polytechnique Federale de Lausanne), with whom it partners in developing certain lines of products. Debiotech is specializing in drug infusion devices,

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using both microfabrication and conventional manufacturing techniques. Debiotech started their microfluidic line of business by developing a silicon micropump under the guidance of Harald Van Lintel, one of the pioneers of microfluidic pump development at University of Twente in the Netherlands. Debiotech has made a number of strategic partnerships, acquisitions, and spinoffs for the development and deployment of drug infusion technology in the medical market. Most notably, in 2007 it signed an agreement with ST Microelectronics for the development and mass fabrication of a MEMS insulin pump, branded Nanopump. The first line of products issued of this collaboration, branded JewelPump, will be applied as a skin patch (awaiting FDA approval in the United States). 5.3.5.2

Technology

The Debiotech technology is based on the development of silicon micromachining for the manufacturing of reciprocating diaphragm pumps. The design used by Debiotech is classic, tracing its roots back to the first published microfluidic pump design by Van Lintel, in 1988 [24]. The technology has been perfected by Debiotech in several design generations, initially under the oversight of Harald Van Lintel himself. The latest generation, branded ChronoJET or Nanopump, is being manufactured and commercialized in cooperation with ST Microelectronics, and is intended for external drug delivery applications. The design includes a silicon membraned actuated by a piezoelectric disk and passive inlet and outlet check valves. The manufacturing involves a three-wafer process (one SOI wafer and two glass wafers), as shown in Figure 5.10. Among the improvements engineered over the time in the micropump design is a high compression ration (allowing the pump to be self-priming), a double limiter concept for the actuation membrane (which assures pumping accuracy irrespective of inlet and outlet pressures or of the fluid viscosity), and a particle filter to stop the occasional particle from entering the pump and the bloodstream. Recently, Debiotech was able to integrate a pressure sensor onto the pump chip [25]. This represents a big step forward in monolythic integration, and it should lead to drug infusion products that are safer (particularly in implantable configurations, where pumping failure conditions such as the occlusion of the channels will be detectable in real time), and have significantly improved accuracy. Another micropump product in development at Debiotech is an implantable model, which has a slightly different construction, based on a four-wafer design: two silicon wafers that incorporate the pumping diaphragm and the check valves, respectively, and two glass wafers which contain the fluidic passages and are bonded to titanium fluidic connectors [26].


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Figure 5.10 Left: Photograph of the Nanopump product from Debiotech, measuring only 16×12×1.9 mm3 (with permission from Debiotech SA / Switzerland) Right: Design of the microfabricated pump based on SOI wafer technology, and incorporating a silicon piezoresistive pressure sensor for the monitoring of pump performance. (Reproduced from [25], with permission from Elsevier.)

Chronoflow is another technology in development at Debiotech, that implements a passive variable flow restrictor using MEMS technology. The idea behind Chronoflow, represented in Figure 5.11 (left), is that a flexible membrane can deform under the action of inlet pressure, thus occluding (or restricting) a number of flow passages and assuring that the flow rate is constant regardless of the pressure difference across the device. The technology allows flow rate control in drug infusion applications, with the flow rate being constant over close to an order of magnitude variation in inlet pressure. The technology can be implemented using the same technologies (e.g., Si/glass micromachining) used in other Debiotech products or using low-cost elastomeric or plastic components. Nanoject represents a third technological brick allowing Debiotech to offer complete drug infusion solution. Nanoject is an array of microneedles fabricated using MEMS technology, allowing intradermal and hypodermic drug delivery as well as interstitial fluid diagnostics. The technology was initially developed at the Royal Institute of Technology in Stockholm, Sweden, and was licensed to Debiotech. The manufacturing involves a specialized DRIE process, capable to create sharp hollow needle arrays on the scale (height and lateral spacing) of 100 microns [27] (Figure 5.11, right). Debiotech has demonstrated an impressive array of technologies that can be combined monolythically to implement advanced drug delivery manipulations, including implantable operation. The devices can be mass produced out of biocompatible materials such as silicon and glass. The current products, aimed at the medical market, have targeted ambient pressure conditions. It would be interesting to see if the same technological bricks

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Figure 5.11 Left: Schematic operation of the MEMS Chronoflow flow limiter concept. (With permission from Debiotech SA / Switzerland) Right: Microneedle array for transdermal drug injection. (Reproduced from [27].)

could be used in applications requiring elevated pressures (as is the case in many industrial cases), by implementing, for example, pressure-balancing mechanisms. It would also be nice to see applications where these bricks become part of larger-scale lab-on-chip implementations, either using a monolythical approach or hybrid integration.

5.3.6 CMOSens Sensor/Electronics Integration: Sensirion 5.3.6.1

History

Sensirion (initially Alpha Sensors) emerged in 1998 as a spinoff from the Swiss Federal Institute of Technology in Zurich. Its technology advantage consists of the capability to integrate on a single chip the sensor part and the associated CMOS signal conditioning, linearization, and temperature compensation electronics (a technology branded CMOSens). It has received a number of awards for its technology (most notably, the 2004 Swiss Economic Award), and it has grown very fast—in 12 years, it has opened several branch offices in addition to its Swiss headquarters, and has established itself solidly in the sensor market, particularly in the area of microfabricated flow measurement devices for the medical, process automation, and automotive industries.


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5.3.6.2 Technology

Sensirion offers several sensor solutions for measuring mass flow (for liquids and gases), differential pressure, temperature, and relative humidity of air and other gases, based on proprietary technology branded CMOSens. CMOSens combines in a monolythic fashion sensor elements based on capacitive, resistive, or thermal (calorimetric) principles, with CMOS electronics for signal digitization, linearization, storage of calibration parameters, and data transmission. The resulting sensor solutions are robust, low-cost, and accurate, due to the complete integration of the technology on a single chip. The flow sensor designed by Sensirion involves a CMOS chip containing heaters and temperature detectors, which attaches to the outside of a fluidic capillary and measures the thermal transport due to flow inside the capillary. The calorimetric principle employed is shown in Figure 5.12. One interesting aspect of the technology is the complete media separation of the sensor part— by acting on the outside of the capillary tube, it allows measurements to be performed on many types of fluids without introducing dead volumes [28]. In addition to the flow rate measurement, further thermal signal processing allows the determination of relative concentration of a mixture of two fluids with different known thermal conductivities based on the combined thermal conductivity. This feature can have applications in many areas, in particular in anesthesiology [29]. The differential pressure sensor manufactured by Sensirion is dynamical (allowing a small amount of fluid passage from one side to the other); in fact, it is a flow meter in disguise: it measures the flow rate, from which (assuming a knowledge of fluid viscosity) the pressure difference can be inferred [30]. The humidity sensor manufactured by Sensirion uses capacitive sensing and a polymer-based dielectric that changes its dielectric properties depending on the amount of adsorbed water, and hence on the environmental humidity. It is integrated with signal conditioning, linearization, and communication electronics resulting in a small and economical sensing solution for many OEM applications such as HVAC. The CMOSens technology can have a huge impact in miniaturizing sensor packages and achieving complete functional integration. It has been applied so far to a number of applications involving thermal and capacitive measurements; however, in principle, it could be extended to many microfluidic applications, including applications where the fluid medium is in contact with the actual sensor parts (which has not yet been realized at Sensirion). By requiring a complete CMOS circuit to be designed around each application, the technology requires a mature sensor architecture that is not subject to many changes or rapid evolution. As microfluidic technology becomes more standardized, and various building blocks are developed, using

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Figure 5.12 Left: Schematic operation of a calorimetric CMOSens flow sensor. A central heater injects flow, through the wall of a capillary tube, into flowing fluid. The heat is advected by the flow, resulting in a temperature difference between the two temperature sensors symetrically located about the heater. (Reproduced from [31].) Right: Photograph of complete flow measurement device, including sensor, electronics, and communication interface. (Courtesy of Sensirion.)

CMOSens or similar technology for more elaborate monolythically integrated applications (chemical sensing, fluidic manipulations, drug delivery, lab-ona-chip) may become feasible. Presently, CMOSens sensors can be integrated with most other microfluidic technologies using hybrid methods. 5.3.7 Integrated Gas Chromatography Microsystem: C2V 5.3.7.1

History

C2V was born out of the strong MEMS and microsystems expertise that developed over the years at University of Twente in the Netherlands and at the associated MESA institute. C2V has its roots in a small start-up company called Twente Microproducts (TMP) that was founded in 1995. TMP used the microfabrication facilities of the University of Twente to commercialize a number of custom products and services, most notably in the area of microsystem design and fabrication. In 2001, TMP was purchased by Alcatel. A year later, a team led by the former managing director of TMP, Job Elders, incorporated Concept to Volume (C2V), and realigned it with the market demands. By 2005 C2V announced the development of microfabricated gas chromatography (microGC) products, and by 2008 it was shipping functional prototypes. It was acquired by Thermo Fisher Scientific in 2009, and finally released its first commercial microGC product (model C2V-200) in 2010. C2V has still maintained a line of business in the area of custom microsystems


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development, and it offers hybrid integration of multiple components through its fluidic platform branded microDELTA. C2V, now a division of Thermo Fisher Scientific, is still being run by Job Elders as the business director. 5.3.7.2 Technology

Gas chromatography (GC) is a complex analytical technique involving the separation of different types of gas molecules in a coated capillary column based on their affinity for the coating material (which is called the “stationary phase”). Different types of column coatings allow gas molecules to be separated based on certain molecular characteristics such as their molecular weight, polarity, and presence of certain functional groups. The gas molecules are advected through the column using a carrier gas, typically hydrogen or helium. The chromatography process involves many individual operations: the precise injection of a small amount of sample gas into the carrier gas flow, separation in the chromatography column, and detection of peaks corresponding to the arrival of different types of molecules in the sample gas mixture at the end of the chromatography column. GC analyses can be further refined by injecting the output of a chromatography column into a mass spectrometer for molecular identification (GC/MS analysis). Sophisticated GC analyses may require an individual separation peak from a first GC column to be injected into a second GC column, a technique called multidimensional GC. A microfabricated GC system requires a number of components to function correcly: preconcentrators, microvalves for injecting the sample into the carrier gas, microfabricated columns coated with different stationary phases, heaters and temperature detectors for controlling column temperature, and sensors for detecting the arrival of different types of molecules. These units need to work very well together, and issues such as dead volumes, thickness of stationary phase, and detector sensitivity and response time can be critical for the operation of the GC system. Temperature stability is also critical for GC operation, as the adsorption/desorption processes responsible for molecular separation in the column are very sensitive to temperature. The issues of microfluidic integration are therefore critical in GC microsystems. The C2V solution for integrating different components is based on a technology platform branded microDELTA, which involves a microfabricated channel plate connecting different units of the GC system together with minimal dead volumes [32]. The platform allows hybrid integration using conventional integrated circuit assembly equipment and technology, allowing components such as sensors and valves to be assembled together. With ICtype bonding techniques (e.g., flip-chip or adhesive), electrical and fluidic

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connections between the channel plate and the components are achieved. The channel plate is fabricated in silicon/glass technology. This solution allows the manufacturing of a very compact and versatile GC cartridge, measuring only 9×5×1.5 cm3 (Figure 5.13), and integrating an impressive numer of individual microfabricated functional units (TCD detectors, valves, heaters, separation column, pressure sensors). This GC cartridge is an easily exchangeable consumable, which reduces support time, and prevents endusers from handling the columns and chips. The cartridge fits in a GC module or channel, whereas the GC instrument can contain 1 to 4 GC channels, each with a different column and detection method. This modular setup allows the ability to monitor a wider spectrum of gas components in the same timeframe of 10 to 60 seconds . The microDELTA technology is one possible solution to the integration problem—it allows each individual component to be manufactured individually, using the most appropriate technologies, as long as certain design parameters are maintained allowing the components to be placed on a channel plate of specific design. The gain in cost and space compared to that afforded by monolythic integration is not impressive; however, the technology is very versatile and allows an integration of technologies that are incompatible from a manufacturing standpoint. One possibility of further improvement would be the integration of CMOS electronics on the microDELTA platform itself (which is a priori possible given that one of materials used for the channel plate is silicon), to allow signal acquisition, processing, digitization, and communication to be integrated. Alternatively, flip-chip IC packages could be used, transforming the channel plate into a silicon circuitboard.

5.4 Conclusions It is apparent from this chapter that much of the innovative development in the area of microfluidics originates near large research institutions with a tradition in microtechnology. There are several reasons for this situation. First, most of the technologies are developed in academic research laboratories initially, and then are commercialized through spinoffs that are often led by scientists and engineers involved in the original research. Second, several research institutions have realized that their reputation as technology innovators depends to a large extent on the commercial success of their technologies, and therefore they are keen to license out the technologies and often help the nascent companies with access to facilities and near-campus real estate. Third, small start-ups rarely have the funds to develop the expensive infrastructure and logistics required for micromanufacturing—the construction of a clean room may cost anywhere between a million and hundreds of millions of


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Figure 5.13 Left: The C2V-200 system from C2V, based on the microDELTA integration platform, and hybridly integrating valves, column, heaters, and TCD detectors. (Used with permission from C2V.) Right: Schematic of the hybrid integration platform, showing the different chips to be assembled, the connecting fluidic channels (4), the elastomeric gaskets (6) for making fluidic connections, gold pads and gold bumps for making electrical connections (8,9), and the adhesive bondings holding the system together (11). (Reproduced from [33].)

U.S. dollars depending on the size, class, facilities, and equipment required. Therefore, access to university clean rooms is essential for such companies in the product development phase. Finally, as technology campuses around such institutions grow, there are interactions between different companies that may lead to synergies and to development of new products leveraging the advantages of multiple technologies.

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[16] ThinXXS marketing brochure: “Perfection in Plastics”, http://www.thinxxs. com/main/download.html. [17] microBUILDER project web site., http://www.microbuilder.org. [18] Project brochure: microBUILDER: An integrated modular service for microfluidics, http://www.thinxxs.com/main/download.html. [19] Vogl, A., “Microbuilder: A Flow-Sensor Module for a Microfluidic Platform,” Europractice MST News, Vol. 6, 2007, p. 26. [20] Smith, R., Sparks, D., Riley, D., and Najafi, N., “A MEMS-Based Coriolis Mass Flow Sensor for Industrial Applications,” IEEE Transactions on Industrial Electronics, Vol. 56, 2009, p. 1066. [21] Sparks, D., Massoud-Ansari, S., and Najafi, N., “Chip-Level Vacuum Packaging of Micromachines Using Nanogetters,” IEEE Transactions on Advanced Packaging, Vol. 26, 2003, p. 277. [22] Sparks, D., “Thin Film Getters: Vacuum Pumps for Microsensors and Actuators,” Vacuum Coating Technology, April Issue, 2010, p. 44. [23] Sparks, D., Smith, R., Cruz, V., Tran, N., Chimbayo, A., Riley, D., and Najafi, N., “Dynamic and Kinematic Viscosity Measurements with a Resonating Microtube,” Sensors and Actuators A: Physical, Vol. 149, 2009, p. 38. [24] Van Lintel, H., Van de Pol, F., and Bouwstra, S., “A Piezoelectric Micropump Based on Micromachining of Silicon,” Sensors and Actuators, Vol. 15, 1988, p. 153. [25] Schneeberger, N., Allendes, R., Bianchi, F., Chappel, E., Conan, C., Gamper, S., and Schlund, M., “Drug Delivery Micropump with Built-in Monitoring,” Procedia Chemistry, Vol. 1, 2009, p. 1339. [26] Maillefer, D., Gamper, S., Frehner, B., Balmer, P., Van Lintel, H., and Renaud, P., “A High-Performance Silicon Micropump for Disposable Drug Delivery Systems,” The 14th IEEE International Conference on Micro Electro Mechanical Systems, 2001. MEMS 2001., 2002, p. 413. [27] Griss, P., and Stemme, G., “Side-Opened out-of-Plane Microneedles for Microfluidic Transdermal Liquid Transfer,” Journal of Microelectromechanical Systems, Vol. 12, 2003, p. 296. [28] Kanne, U., “Digital Cmos Sensor Chips for Media-Isolated Liquid Flow Sensing.” Medical Device Technology, Vol. 14, 2003, p. 16. [29] Kanne, U., “Digital Sensor Chips for Agent Dosing and Metering,” Medical Device Technology, Vol. 16, 2005, p. 17. [30] Sensirion Application Note for SDP600 and SDP1000 Series, http://www. sensirion.com/en/04_differential_pressure_sensors/00_ differential_pressure_sensors.htm. [31] Kanne, U., and Sauvain, C., “Digital Flow Sensors: Reaching New Levels,” Medical Device Technology, Vol. 17, 2006, p. 12. [32] Van Weerden, H., Burger, G., and Elders, J., “Fast Micro-GC Capabilities Based on a

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Microintegration Technology Platform,” American Laboratory, Vol. 40, 2008, p. 14. [33] Burger, G., Vis, J., and Van Weerden, H., “Method for Building a Device Having Fluidic and Electrical Functions (U.S. Patent App. 20,080/250,633),” , 2008.

About the Author Dan E. Angelescu is a professor in the Electronics Systems Department at Université Paris Est (ESIEE Paris). His research is focused on the development of integrated microfluidic systems for performing fundamental science experiments as well as for specific industrial applications. Prior to his current academic position, he was a senior research scientist at the Schlumberger research laboratory in Boston, where he developed microfluidic solutions for sensing in extreme environments. Dan E. Angelescu maintains a consulting activity in the area of microsystem design. He holds a B.S. in Physics from the California Institute of Technology, and a Ph.D. in Physics from Princeton University.

249

Index capillary electrophoresis, 122 centrifugal pump, 77 CFD, 214 channel, circular, 157 channel, rectangular, 159 chaotic advection, 95 check valve, 72 chemical measurement, 120 chemical requirements, 47 chemical resistance, 47 chemiluminescence, 115 circular channel, 157 clean room, 194 CMOS integration, 241 contact angle, 175 controlled-release valve, 73 Coriolis flowmeter, 124, 237 cost, 48 Coventor, 207 CVD, 23

3-D structures, 28 absorption spectroscopy, 118 ACEO, 180 ACEO pump, 90 active mixing, 97 active valve, 58 actuation, electromagnetic, 58 actuation, electrostatic, 67 actuation, piezoelectric, 61 actuation, pneumatic, 63 additive processes, 21 adhesive bonding, 16, 35 ALD, 23 alignment marks, 209 annealing, 31 anodic bonding, 33 atomic layer deposition, 23 bonding, 7, 12, 16, 33–35 bonding, adhesive, 16, 35 bonding, anodic, 33 bonding, diffusion, 34 bonding, eutectic, 34 bonding, metal, 12 bonding, PDMS, 7 bonding, plastic, 16 brazing, 12 Brownian motion, 141 burst valve, 73

dead volume, 161, 191 Debiotech SA, 237 density sensor, 124, 237 detection, 114 dicing marks, 209 dielectrophoresis, 179 dielectrophoretic, 108 diffusion, 141 diffusion bonding, 34 digital microfluidics, 99 direct bonding, 34

C2V, 242 CAD tools, 206 251

252


doping, 32 DRIE, 27 droplet, 99 droplet generation, EWOD, 105 droplet generator, 101, 200 droplet generator, dielectrophoretic, 105 droplet generator, flow-focusing, 101 droplet generator, microchannel emulsification, 104 droplet generator, T-junction, 101 droplet generator, topological, 104, 201 droplet microfluidics, 230 droplet mixing, 107 droplet, coalescence, 110 droplet, manipulation, 108 droplet, splitting, 110 drug infusion, 239 dry etching, 27 e-beam writing, 21 EDM, 9 electric interconnects, 42 electrical analogy, 166, 191, 211 electro-osmosis, 180 electro-osmotic pump, 88 electrochemical, 122 electrohydrodynamic pump, 88, 90 electrokinetic, 98 electromagnetic actuation, 58 electromagnetic force, 151 electromagnetic valve, 58 electrophoresis, 178 electroplating, 24 electrostatic actuation, 67 electrostatic pump, 82 electrostatic valve, 67 electrowetting, 181 embossing, 15 emulsion, 101, 200 entry/exit effects, 159 etching, 11, 25, 27 eutectic bonding, 34 evaporation, 22 EWOD, 105 fabrication process, 196 FIB, 31 flow control, 38 flow focusing, 96, 101

flow sensing, 39 flowmeter, anemometry, 125 flowmeter, Coriolis, 124, 237 flowmeter, themal, 241 flowmeter, time of flight, 126 fluid dynamics, 143 fluid forces, 147 fluidic interconnects, 41 Fluidigm, 225 fluorescence, 115 focused ion beam, 31 formative processes, 31 foundry, 194 functional approach, 56, 187 functional blocks, 38, 57 functional diagram, 36 gas chromatography, 243 glass, 17 gravitational force, 151 Hadamard-Rybczynski, 165 Hele-Shaw geometry, 155 high-speed milling, 8 hybrid integration, 188, 243 hydrodynamic capacitance, 167 hydrodynamic resistance, 166 hydrodynamic valve, 74 IFC, 225 injection molding, 14, 232 insulin pump, 238 integration, 43, 222 integration, hybrid, 188 integration, monolythic, 188 interconnects, 41 interconnects, electric, 42 interconnects, fluidic, 41, 189 ion milling, 31 ISSYS, 235 Kirchhoff’s laws, 168 laser ablation, 11 laser writing, 21 latching valve, 68 liftoff, 25 LIGA, 10 lithography, 19 lithography layer, 208

253

Index

magnetohydrodynamic pump, 88, 90 manipulation, droplet, 108 manipulators, 40 manufacturing, 194 Marangoni stresses, 176 market, 223 mask design, 205 mass conservation, 144 materials choices, 193 measurement, 114 measurement, chemical, 120 measurement, optical, 115 measurement, physical, 123 MEMS, 17 MEMSNAS, 28 meshing, 215 metal bonding, 12 Metal fabrication, 7 micro-EDM, 9 microelectronics, 2 microfluidics, 2 microfluidics market, 223 microfluidics, multiphase, 99 microfluidics, single-phase, 57 micromachining, 8 micropump, 197 microtube, 236 mixing, active, 97 mixing, chaotic, 98 mixing, chaotic advection, 95 mixing, droplet, 107 mixing, electrokinetic, 98 mixing, flow focusing, 96 mixing, flow lamination, 94 mixing, passive, 93 mixing, single-phase, 91 mixing, T-junction, 98 mLSI, 225 molding, 3 momentum conservation, 145 monolythic integration, 188 multiphase microfluidics, 99 nanoimprint, 20 Navier-Stokes, 151 non-Newtonian fluid, 148 optical measurement, 115 optical requirements, 45

optimization, 205, 210 oxidation, 32 paper fabrication, 17 parabolic profile, 156 passive mixing, 93 passive valves, 71 patents, 224 PCR, 188 PDMS, 3, 86, 117, 197, 229 PDMS bonding, 7 PE-CVD, 23 peristaltic pump, 85, 87, 197, 212 phase separation, 113 photoetching, 11 photolithography, 20 physical requirements, 45 physical, measurement, 123 piezoelectric actuation, 61 piezoelectric pump, 81, 83, 84, 87, 233, 238 piezoelectric valve, 61 planarization, 30 plastic bonding, 16 plastic fabrication, 14 plating, 24 pneumatic actuation, 63 pneumatic valve, 63 Poiseuille, 160 polishing, 30 powder blasting, 28 pressure gradient, 147 pressure requirements, 47 pressure sensing, 40 pressure-driven, 76 process chart, 200, 202 process, fabrication, 196 processes, additive, 21 processes, formative, 31 processes, lithographic, 19 processes, subtractive, 25 pump, ACEO, 90 pump, centrifugal, 77 pump, electro-osmotic, 88 pump, electrohydrodynamic, 88, 90 pump, electrostatic, 82 pump, magnetohydrodynamic, 88, 90 pump, merit criteria, 79 pump, peristaltic, 85, 87, 197, 212

254


pump, piezoelectric, 81, 83, 84, 87, 233, 238 pump, reciprocating, 77 pump, syringe, 75 pumping techniques, 75

subtractive processes, 25 surface tension, 172 surfactant, 176 syringe pump, 75 system design, 205

RainDance Technologies, 229 reciprocating pump, 77 recirculating flow, 163 rectangular channel, 159 requirements, chemical, 47 requirements, optical, 45 requirements, physical, 45 requirements, pressure, 47 requirements, temperature, 46 requirements, wetting, 46 resistance, hydrodynamic, 166 Reynolds number, 152 rheology, 124, 150 RIE, 27

T-junction, 98, 101 Taylor dispersion, 160 technology choice, 186 technology examples, 225 technology selection, 35 temperature requirements, 46 temperature sensing, 39 thermal bonding, 17 thermal diffusion, 141 thermal treatment, 39 ThinXXS Gmbh, 232 time constant, 170 transient, 170

sandblasting, 28 sensing, 114 Sensirion, 240 sensor, density, 124 sensor, viscosity, 124 shadow-mask, 20 silicon, 17 silicon on insulator, SOI, 18, 81 silicon tube, 235 silicon/glass fabrication, 17 single-phase microfluidics, 57 single-phase mixing, 91 soap, 176 Soft Lithography, 3 soft lithography, 197 solvent bonding, 16 spin coating, 21 sputtering, 22 Stokes flow, 144, 154

ultrasonic machining, 30 valve, active, 58 valve, burst, 73 valve, check, 72 valve, electromagnetic, 58 valve, electrostatic, 67 valve, hydrodynamic, 74 valve, latching, 68 valve, piezoelectric, 61 valve, pneumatic, 63 valves, passive, 71 viscosity, 148, 164 viscosity sensor, 124, 237 viscous drag, 164 viscous heating, 171 welding, 13 wet etching, 25 wetting, 174 wetting requirements, 46

Recent Titles in the Artech House Integrated Microsystems Series Acoustic Wave and Electromechanical Resonators: Concept to Key Applications, Humberto Campanella Adaptive Cooling of Integrated Circuits Using Digital Microfluidics, Philip Y. Paik, Krishnendu Chakrabarty, and Vamsee K. Pamula Fundamentals and Applications of Microfluidics, Second Edition, Nam-Trung Nguyen and Steven T. Wereley Highly Integrated Microfluidics Design, Dan E. Angelescu Integrated Interconnect Technologies for 3D Nanoelectronic Systems, Muhannad S. Bakir and James D. Meindl, editors Introduction to Microelectromechanical (MEM) Microwave Systems, Héctor J. De Los Santos An Introduction to Microelectromechanical Systems Engineering, Nadim Maluf Lab-on-a-Chip, Techniques, Circuits, and Biomedical Applications, Yehya H. Ghallab and Wael Badawy MEMS Mechanical Sensors, Stephen Beeby et al. Micro and Nano Manipulations for Biomedical Applications, Tachung C. YihIlie Talpasanu Microfabrication for Microfluidics, Sang-Joon John Lee and Narayan Sundararajan Microfluidics for Biotechnology, Second Edition, Jean Berthier and Pascal Silberzan Organic and Inorganic Nanostructures, Alexei Nabok Post-Processing Techniques for Integrated MEMS, Sherif Sedky ( Pressure-Driven Microfluidics, Václav Tesar RFID-Enabled Sensor Design and Applicatons, Amin Rida, Li Yang, and Manos Tentzeris RF MEMS Circuit Design for Wireless Communications, Héctor J. De Los Santos Wafer-Level Testing and Test During Burn-in for Integrated Circuits, Sudarshan Bahukudumbi Krishnendu Chakrabarty Wireless Sensor Network, Nirupama Bulusu and Sanjay Jha

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Highly Integrated Microfluidics Design

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Design of Analog CMOS Integrated Circuits

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