OPTOTHERMAL SYSTEMS FOR CELLULAR … · interesting experiences, ... 2.2 Sensors ... therapy, transplant of the affected organ, or use of medical supportive devices [8]

OPTOTHERMAL SYSTEMS FOR CELLULAR MANIPULATION AND TISSUE

CULTURE

A THESIS SUBMITTED TO THE GRADUATE DIVISION OF THE UNIVERSITY OF HAWAI‘I AT MĀNOA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS

FOR THE DEGREE OF

MASTER OF SCIENCE

IN

ELECTRICAL ENGINEERING

MAY 2012

By

Kelly Ann S. Ishii

Thesis Committee:

Aaron Ohta, Chairperson Olga Boric-Lubecke

Victor Lubecke

Keywords: MEMS, microfluidics, biomedical, microrobotics

ii

ACKNOWLEDGMENTS

I wish to express my sincerest gratitude to the numerous individuals who have

played an integral part in my graduate studies and in the culmination of this thesis.

Firstly, I would like to thank my advisor, Dr. Aaron Ohta, for mentorship and

invaluable guidance throughout my time here, for the opportunity to take part in some

interesting experiences, and finally for considerable support in editing and preparation of

my thesis and defense.

I would also like to thank Dr. Olga Boric-Lubecke and Dr. Victor Lubecke for

taking time out of their busy schedules to participate in my committee, and appreciate

their helpful input toward further improvement of this manuscript.

I am also grateful for my laboratory colleagues, who have provided valuable

discussions and suggestions. In particular, I would like to thank Wenqi Hu. As the lead

for the microrobotics projects, his contributions outnumber my own, and include the

development of the theory with initial characterizations and simulations, which he has

kindly allowed me to use here, as well as handling a good part of the microfabrication

used in competition.

The completion of this thesis is also owing to Kory Kurokawa, for lending his

technical expertise and being more than helpful during laboratory work, Dr. Hongwei Li

and Yujia Li for supplying MDCK II cells, and Dr. Shiv Sharma and Dr. Ava Dykes for

supplying HeLa cells.

Finally, I would like to thank my parents, extended family, and friends for their

encouragement and endless support along the way.

iii

ABSTRACT

Two novel optothermal microfluidic systems capable of facilitating tissue cell

culture were developed and tested. Both systems feature optical-based heating that is

used to trigger thermal effects.

The first system involves a mechanism for cell isolation and culture. Cells are

suspended in a thermoresponsive hydrogel solution, and optical patterns are utilized to

heat the solution, producing localized hydrogel formation around cells of interest. The

hydrogel traps only the desired cells in place, while also serving as a biocompatible

scaffold for supporting the cultivation of cells in 3D. This is demonstrated with the

trapping of MDCK II and HeLa cells. The light intensity from the optically induced

hydrogel formation does not significantly affect cell viability.

The second system utilizes optically actuated bubbles in oil as microrobots.

Simulations of the thermocapillary fluid flow within the oil phase are used to illustrate

the mechanisms driving the bubble actuation. Parallel manipulation of sub-millimeter

objects including glass beads and hydrogel beads was demonstrated. These capabilities

show the potential for using the bubble microrobots in a broader range of biomedical or

other microassembly applications, including the assembly of cell-laden scaffolds for

tissue culture.

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TABLE OF CONTENTS

ACKNOWLEDGMENTS .................................................................................................. ii ABSTRACT ....................................................................................................................... iii

LIST OF FIGURES ........................................................................................................... vi LIST OF ABBREVIATIONS ........................................................................................... vii

CHAPTER 1. INTRODUCTION ........................................................................................1 1.1 Background ................................................................................................................2 1.2 Organization ...............................................................................................................5

CHAPTER 2. OPTOTHERMAL DEVICES .......................................................................7 2.1 Introduction ................................................................................................................7 2.2 Sensors .......................................................................................................................7 2.3 Actuators ....................................................................................................................7 2.4 Fluid and Particle Manipulation .................................................................................8 2.5 Tuning ......................................................................................................................10 2.6 Conclusion ...............................................................................................................10

CHAPTER 3. CELL TRAPPING AND CULTURE .........................................................11 3.1 Introduction ..............................................................................................................11

3.1.1 Background .......................................................................................................12 3.2 Experimental Setup ..................................................................................................14

3.2.1 Culture Device ..................................................................................................15 3.2.2 Culture Protocol ...............................................................................................16

3.3 Materials and Methods .............................................................................................17 3.3.1 Thermoresponsive Polymer ..............................................................................17 3.3.2 Cells ..................................................................................................................17

3.4 Results ......................................................................................................................18 3.4.1 Simulation .........................................................................................................18 3.4.2 Substrate Characterization and Selection ........................................................19 3.4.3 Mammalian Cells ..............................................................................................21

3.5 Discussion ................................................................................................................25 3.5.1 Evaluation of Trapping on Viability .................................................................25 3.5.2 Considerations for Long-term Culture .............................................................26

3.6 Conclusion ...............................................................................................................27 CHAPTER 4. BUBBLE MICROROBOTS .......................................................................28

4.1 Introduction ..............................................................................................................28 4.1.1 Background .......................................................................................................28

4.2 Experimental Setup and Control System .................................................................30 4.2.1 Optical Patterns ................................................................................................31 4.2.2 Microrobot Chamber ........................................................................................31 4.2.3 Control Software ...............................................................................................31

4.3 Actuation Mechanisms and Theory .........................................................................32 4.4 Materials and Methods .............................................................................................35

4.4.1 Microfluidic Holding Chamber .........................................................................35

v

4.4.2 Bubble Microrobot ............................................................................................36 4.4.3 Micro-objects ....................................................................................................37

4.5 Results ......................................................................................................................38 4.5.1 Force on a Glass Bead ......................................................................................38 4.5.2 Microassembly of SU-8 Triangles ....................................................................39 4.5.3 Parallel Actuation and Microassembly ............................................................40

4.6 Discussion ................................................................................................................43 4.6.1 Variations due to Operating Environment ........................................................43 4.6.2 Considerations for Further Development .........................................................43

4.7 Conclusion ...............................................................................................................45

CHAPTER 5. CONCLUSION ..........................................................................................47 APPENDIX ........................................................................................................................49

A1 Cell Handling Protocol for Trapping/Culture Device ..............................................49 A2 Viability Stain Protocol for Trapping/Culture Device ............................................50

REFERENCES ..................................................................................................................51

vi

LIST OF FIGURES Fig. 3.1 Example cell isolation procedure .........................................................................13

Fig. 3.2 Culture device experimental setup .......................................................................15

Fig. 3.3 Culture device schematic .....................................................................................16

Fig. 3.4 Cell trapping and culture process ........................................................................17

Fig. 3.5 Simulated temperature distribution .....................................................................19

Fig. 3.6 Hydrogel area vs. exposure time .........................................................................20

Fig. 3.7 Hydrogel area vs. laser power .............................................................................21

Fig. 3.8 Single and multiple trapping ................................................................................23

Fig. 3.9 MDCK II viability ...............................................................................................24

Fig. 3.10 HeLa viability ....................................................................................................25

Fig. 4.1 Microrobot experimental setup ............................................................................30

Fig. 4.2 Simulated thermocapillary flow ..........................................................................34

Fig. 4.3 Microrobot chamber designs ...............................................................................35

Fig. 4.4 Force on glass bead vs. bubble radius .................................................................39

Fig. 4.5 Parallel manipulation of six microrobots .............................................................40

Fig. 4.6 Cooperative manipulation of agarose beads between two microrobots ..............41

Fig. 4.7 Cooperative manipulation of glass bead between three microrobots ..................42

vii

LIST OF ABBREVIATIONS

A-Si Amorphous silicon CDC Centers for Disease Control COPD Chronic obstructive pulmonary disease DEP Dielectrophoresis DI Deionized DMEM Dulbecco’s Modified Eagle Medium ECM EthD-1

Extracellular matrix Ethidium homodimer-1

HBSS Hank’s Balanced Salt Solution ICRA International Conference on Robotics and Automation IEEE Institute of Electrical and Electronics Engineers IPA Isopropyl alcohol ITO Indium tin oxide LCST Lower critical solution temperature MDCK Mandin Darby Canine Kidney NIST National Institute of Science and Technology PBS Phosphate-buffered saline PDMS PhC

Polydimethylsiloxane Photonic crystal

PNIPAAm P(VDF-TrFE)

Poly(N-isopropylacrylamide) Poly(vinylidene trifluoroethylene)

OET TGF

Optoelectronic tweezers Temperature gradient focusing

1

CHAPTER 1

INTRODUCTION

According to preliminary statistical data compiled by the Centers for Disease

Control (CDC), the national leading causes of death between 2008-2010 were 1) heart

disease; 2) cancer; 3) chronic lower respiratory diseases, a category that includes chronic

obstructive pulmonary disease (COPD) and asthma; and 4) cerebrovascular diseases,

most commonly associated with stroke [1], [2]. Chronic diseases such as these do not

usually resolve spontaneously and are seldom completely cured in response to treatments

[3]. In the United States, these diseases account for approximately every seven out of ten

deaths, and nearly half of the adult population is currently living with at least one of these

conditions. Even those who do not experience these diseases firsthand can be indirectly

affected. More than 75% of the national health care expenditures are attributed to chronic

diseases. For heart disease and cancer alone, the estimated cost of health care services,

medications, and loss of worker productivity in 2010 totaled $372.7 billion [4], [5].

Diabetes, another highly prevalent chronic disease in the United States and other

countries, ranks seventh, but estimates project that the pervasiveness of this disease will

continue to increase [6]. Most diabetes cases are not immediately life threatening, but it

can still have a significant impact on overall health. The presence of one disease may

contribute to the development of another; for example, heart disease and diabetes elevate

the risk of stroke [7]. Complications that arise as a result of disease can require extensive

treatment, potentially cause disability, or otherwise reduce the quality of life. Current

treatment options for chronic diseases are few, usually restricted to long-term drug

therapy, transplant of the affected organ, or use of medical supportive devices [8]. There

are shortcomings associated with each option, and only organ transplant offers more than

simply managing the disease.

Regenerative medicine a broad term encompassing concepts like cellular therapy,

tissue engineering, biomedical engineering of artificial organs, and gene therapy, is a

promising solution to a variety of medical conditions [8]. The first cell-based therapy,

bone marrow transplant, dates back more than five decades [9]. Since then, cellular

therapies have been explored for several medical conditions [10], [11] including heart

2

disease, cancer, chronic lower respiratory disease, cerebrovascular disease, and diabetes

[12]-[20].

Unlike cellular therapy, tissue engineering is a newer, developing field that has

not seen as much clinical use. Tissue engineering is a cross-disciplinary effort that aims

to restore the body's normal function by replacing diseased, injured, or defective parts.

This can involve the creation of new functional tissues or organs for transplant by

culturing cells into 3D structures or by encouraging endogenous self-repair with the

implantation of scaffolds that provide a template for cell growth.

In addition to regenerative medicine, another area in biomedicine that stands to

gain from improved culture techniques is cell-based drug assays. The time from drug

discovery to Food and Drug Administration approval is usually 10 to 15 years [21],

costing from $800 million to $1 billion [22]. Requisite testing, although necessary for

safety and ultimately beneficial for the consumer, can lengthen the process of identifying

suitable treatments. Improvements to existing cell-based assays have the potential

transform the way drug testing is conducted. Cells cultivated in 3D environments are

more representative of their in vivo counterparts and thus are more predictive of actual

drug response [23]. As the ability to culture cells in 3D improves, the practice of testing

on live subjects could be reduced, mitigating risk of testing revolutionary, but potentially

dangerous drugs, yet encouraging them to be explored. A more distant possibility is

treatment tailored to individual patients, with tests built off a culture of their own cells.

1.1 Background

The majority of mammalian cells are adherent, meaning that they require a

surface to grow on. There are some exceptions, such as blood cells, which are naturally

suspended in fluid. From this point on, discussions of cell culture will pertain to adherent

cells. Cells used in transplants originate from one of three sources [24]. Autologous cells

are obtained directly from the intended recipient of a transplant. This type of cell is the

most compatible with the immune system but is the least economical in terms of

production. The use of autologous cells also may not be feasible due the inability to

generate a sufficient amount of cells within time constraints. Allogeneic cells are taken

from human sources other than the patient. Tissues derived from these cells have the

3

potential to be mass-produced, but come with a greater chance of transplant rejection.

Xenogeneic cells are acquired from non-human sources. Of the three, these would be the

most abundant; however, extra effort would be devoted to ensuring safety and immune

acceptance.

Conventionally, cells maintained for routine laboratory use are kept in 2D

cultures, which consist of cells grown to monolayers on surfaces that promote cell

adhesion. 2D cultures are cost effective and low maintenance. Care of 2D cultures

includes regular feeding, or replenishment of the culture’s nutrient supply by replacing

old media with fresh media, and passaging, or the harvesting, splitting, and re-seeding

cells to avoid overcrowding on the culture plate. Feeding and passaging are the only

manual cell handling required in this process, while equipment such as incubators, and

media formulations handle the other requirements for cell growth.

Histotypic and organotypic cultures describe cultures in which the constituent

cells have been organized into 3D structures instead of 2D, that are reminiscent of in vivo

tissues and organs, respectively [25]. Even if an application does not require a tissue,

such as drug assays, recent evidence suggests that 3D cultures yield better representations

of cells in vivo compared to those raised in 2D cultures, as 3D cultures preserve some

semblance of the natural cellular environment that fosters important cell-cell and cell-

extracellular matrix (ECM) interactions [26]. For example, cancerous cells grown in 3D

cultures are more resistant to drugs versus the same cells grown in 2D cultures [27], [28].

However, 3D cultures are understandably more difficult to establish, control, and

maintain.

In general, 3D culture has largely focused on two methods [24]. In one method,

an acellular scaffold mimicking the native ECM is implanted into the body where it is

populated by autologous cells. The scaffold serves as a guide for cell migration and gives

the tissue formation mechanical strength. In the other method, cells are seeded onto a

scaffold and raised in vitro. Seeding techniques have been developed where a cell

suspension is deposited onto porous biomaterials [21], [28], [29] and other prefabricated

microstructures [30], [31] or directly mixed into scaffold material before the architecture

is defined [32]. More recently—beginning in the late 1990s—scaffoldless methods have

emerged, where cells are grown in sheets are layered on top of each other [33], [34].

4

Each method has its own advantages and disadvantages. Both scaffold-based

methods allow the definition of complex supporting architectures to guide tissue

development. However, the selection of an appropriate scaffold can have a significant

impact not only on cell viability, but also behavior of the culture as a tissue. A variety of

scaffold materials have been documented, derived from both natural and synthetic

sources [35], [36]. At the very least, scaffold materials should be biocompatible and

biodegradable [24]. A scaffold’s internal architecture, consisting of a network of pores,

should facilitate cell migration, adhesion, and mass transport of gases, nutrients, and

waste in and out of the scaffold. The acellular scaffold method eliminates the need to

cultivate tissues in vitro, but with less control over the development process and

localization of the cells, the chance of failure is greater. Reliance on autologous cells also

comes with particular advantages and disadvantages as discussed earlier. Additionally,

the method is only appropriate for applications in regenerative medicine, and not in drug

discovery.

Scaffoldless methods use ECM secreted by the cells themselves, keeping the

quality of the matrix more consistent with natural conditions, with no need of engineering

an artificial ECM. The complex structures and compositions of naturally occurring ECM

are tissue-specific, and have been observed to dynamically respond to fluctuations in cell

metabolic activity, mechanical demands, and other influences from the surrounding

environment [24], [37]. Recent studies suggest the responsibility of the ECM in

regulating key cell processes (e.g., migration, attachment, and differentiation) is the result

of a combination of structural, mechanical, and chemical interactions with the resident

cells. The implantation of vascular grafts with synthetic matrices revealed problems with

in vivo performance including low patency rates (ability of the graft to remain open), the

possibility of reaction resulting in inflammation, and infection due to bacterial

colonization [33]. Using scaffoldless grafts with ECM produced by the constituent cells

is theoretically advantageous by reducing these problems, in addition to giving the body

the ability to renew and remodel the ECM like it would normally. A downside to this

technique is the long cultivation time that is sometimes necessary to produce complex

layered structures (over 3 months).

5

In all of these techniques, mass transport is a major obstacle in the creation of

tissues and organs. In the absence of vasculature, passive diffusion from the media-

scaffold interface is the mode of transport. Diffusion degrades as the distance from the

surface increases, thus cells near the core of a 3D culture can become necrotic due to the

lack of nutrients [34], [37]. Microfluidic cell culture systems have been developed that

address nutrient and waste delivery through continuous perfusion in a controlled, sterile

environment [21], [23]. Generally restricted to sub-millimeter thicknesses, cultures in

microfluidic devices maintain a relatively high surface-area-to-mass ratio, also aiding in

diffusion. Although these features are beneficial for cell viability, these devices may

interfere with cell-cell signaling, and the issue remains of how to transition to thicker

tissue constructs and organs [21].

Additional considerations arise due to the need to preserve the functionality of the

tissue being cultured; however, a lot less is understood about the exact role of the ECM in

this process [37]. Normally, tissues consist of a heterogeneous mix of cells, thus the

ability to specify the spatial organization of cells is desirable. Cell-cell interaction across

different cell types plays a role in cellular expression. Seeding techniques where cells can

be patterned are favorable in this regard.

The objective of this research is to develop systems that improve the

controllability of tissue culture and related biomedical laboratory protocols. The first

method presented uses a microfluidic device loaded with a cell suspension in scaffold

material for 3D cell patterning. Unlike other patterning techniques, this method enables

selection of constituent cells in a 3D culture, down to individual cells. The second

method presented also based on a microfluidic platform uses a microrobotic system to

carry out micromanipulation tasks. As an application in cell culture, cell-laden scaffolds

are assembled to construct the body of a 3D culture. Both methods aim to improve the

ability to create heterogeneous cultures.

1.2 Organization

This paper is organized as follows. Chapter 2 highlights other instances of

optothermal effects being used in MEMS devices.

6

Chapter 3 describes an optothermal microfluidic device and corresponding control

system with application in 3D tissue culture. The optothermal effects are used to induce

the selective trapping of desired cells in a hydrogel scaffold that can subsequently be used

to culture the cells in 3D. Different absorbent substrates are investigated, and one is

selected for the use in experiments on cells. Single cell trapping and multiple cell

trapping are demonstrated. The parallel trapping of cells within separate scaffolds is also

shown. Performing preliminary viability test on two types of mammalian cells assesses

the suitability of the technique.

A similar optothermal design that supports the actuation of bubble microrobots is

presented in Chapter 4. The system, which consists of air bubbles actuated in oil,

incorporates the same absorbent substrate and control system as the cell-trapping device.

The microrobots are capable of exerting sufficient force to move other micro-objects and

thus can be used as tools for micromanipulation. Results demonstrating microassembly

and cooperation between multiple microrobots also presented.

Chapter 5 summarizes the findings of each work and discusses the potential and

future improvements to each of the techniques.

7

CHAPTER 2

OPTOTHERMAL DEVICES

2.1 Introduction

Optothermal effects have been explored for the operation of several types of

MEMS devices. Power delivery realized through optical pathways rather than traditional

electrical ones can be considered for situations where scaling down makes fabrication

increasingly challenging, contactless control is preferable to wired connections, or when

reconfigurabilty is desired. The objective of this chapter is to explore the breadth of

optothermal applications in engineering outside the research areas presented in

subsequent chapters.

2.2 Sensors

Pyroelectric materials are anisotropic materials that exhibit a temperature-

dependent net polarization. Infrared and millimeter wavelength sensors based on

pyroelectric response have been developed, both in research and commercial

applications.

Zirkl et al. presented a flexible optothermal infrared sensor, which included a

temperature sensitive polymer with pyroelectric properties coupled with an organic field

effect transistor [38]. A ferroelectric poly(vinylidene trifluoroethylene) (P(VDF-TrFE))

copolymer thin film between two aluminum electrodes served as the sensor element. The

bottom electrode is shared with the transistor, serving as the gate. Shining an infrared

laser diode on the top electrode causes the sensor to act as a switch for the attached

transistor.

2.3 Actuators

Thermally driven actuators with electrothermal control have been demonstrated in

the past. Some limitations of electrothermal actuators in MEMS devices are the need for

an external power supply that is wired to the resistive elements, which can in turn lead to

increased device size, packaging difficulty, and undesirable effects such as interference

8

from the current used to cause Joule heating [39]. In response, thermal actuators based on

optothermal expansion of the actuator components have been pursued as an alternative.

Zhang et al. developed bi-directional optothermal microactuators. [40] Tethered

at one end, two thin, parallel actuator beams stretch out from the base and form a

rectangular joint at the untethered ends. Asymmetrical heating by laser irradiation (650

nm) on one of the beams produces deflection toward the unheated beam due to thermal

expansion of the material in the irradiated spot. Black polypropylene was used to

fabricate the microactuators, since it was compatible with excimer laser micromachining

while also offering desirable properties like low thermal conductivity and a high thermal

expansion coefficient. Larger deflections at lower powers could be achieved with

optothermal control as compared to some electrothermal actuators.

Optically controlled polysilicon thermal actuators were investigated as a potential

driving mechanism for microrobots by Szabo et al. [41]. Chevron-style actuators were

heated by irradiation with a 660 nm laser, which causes thermal expansion of the actuator

beams. During expansion, the two halves of the actuator to press together and buckle at

the joint, inducing a linear deflection. The intensity of the laser spot correlates to the

internal energy produced. A set of thermal actuators that orient downwards when

actuated was also included to give the microrobot two degrees of freedom.

2.4 Fluid and Particle Manipulation

Optothermal effects generated in a liquid can be exploited for non-contact

manipulation. Various researchers have demonstrated ways in which optothermal

manipulation can be useful in a microfluidic environment.

Weinert et al. used optical irradiation with a pulsed 1450-nm infrared laser to

generate dynamic flow patterns in a thin layer of water (5-10 µm) sandwiched between

two glass layers [42]. Moving the laser spot initiated a pumping action within the fluid

layer. Temporarily heating a spot produces a change in the density of the water, causing

divergent flows in response to the disequilibrium. As the laser spot moves, the water at its

front expands, while that in its wake contracts. Heating also lowers the viscosity of the

water, thus increasing the flow velocity in the heated region. The combined effects

induce a net flow opposite to the direction that the spot is traveling, creating a laser-

9

driven pump. This mechanism can be extended to the transportation of biomolecules,

which was demonstrated in the creation of a dilution series in agarose gel.

Duhr and Braun used thermophoresis, or the migration of molecules in the

presence of a temperature gradient, for the trapping of DNA in a microchannel [43].

Thermophoresis—also referred to as thermal diffusion or the Soret effect—induces the

movement of molecules from the hotter side to the cooler side of a temperature gradient.

To trap the particles, the thermophoretic drift is countered by fluid flow along the same

axis but in the opposite direction. Localized heating was established by shining a 1480-

nm infrared laser on the channel. Switching from a spot to a radial pattern of

concentrically converging rings creates a centrally located accumulation.

Thermophoresis traps can also be created when thermophoretic flow is directed in

the axis perpendicular to a bidirectional convective flow [42]. The displacement caused

by thermophoresis deposits the molecules into one of the two flow streams. If the

uniaxial temperature gradient in chamber is oriented such that the hot side is on the lower

layer and the cool side is at the top layer, the molecules should respond by moving

toward the top. Bidirectional flow that circulates away from the center of a chamber

along the lower surface and back toward the center along the top surface will cause the

molecules that have drifted to the surface to accumulate at the center of the chamber.

Liu et al. demonstrated optothermal levitation and manipulation of small

microparticles (6-µm-diameter polystyrene beads) in water using a 1550-nm laser [44].

They theorize that optical heating within a fluidic environment can induce convection

fluid flow affecting particles at longer ranges, while radiation forces from the laser beam

transfer energy to the particles, causing attraction in the area immediately surrounding the

laser focal point.

As a method of charged analyte concentration and transportation, important steps

in diagnostic analysis procedures handling small or dilute sample volumes, Akbari et al.

proposed the use of optothermal heating to perform temperature gradient focusing (TGF)

[45]. During TGF, both a potential and a temperature gradient are applied across the

length of a microchannel. Analytes are concentrated at the point in the channel where the

bulk flow of the analytes reaches equilibrium with the particle velocity driven by the

temperature-dependent electrophoresis. Heat is typically delivered through some form of

10

resistive heating element comprising or embedded in a substrate. The method proposed

by the authors involves local heating provided by a commercial projector and focusing

optics, yielding the advantages of flexibility and a simplified system over integrated

heaters. To aid in the conversion of light to heat, black electrical tape was affixed to the

microchannel.

2.5 Tuning

Reversible optothermal tuning of photonic crystal (PhC) nanocavities using liquid

crystal infill was demonstrated by Dundar et al. [46]. The temperature dependence of the

refractive index can be applied as a mechanism for tuning, which may be necessary to

provide functionality or to correct for errors arising from fabrication. Localized heat can

be provided through the inclusion of resistive microheaters; however, optothermal heat

generated by a laser is advantageous because it is a contactless and reconfigurable

method of localized heating that does not require much sample processing. Methods for

improving the tuning of the PhC devices include adding an extra film of absorptive

material, thermally insulating the devices from the rest of the chip, increasing absorption

of the light into the semiconductor, or adding liquid into the PhC cavities.

2.6 Conclusion

In this chapter, applications of optothermal effects were discussed, including

sensors, actuators, fluid and particle manipulators, and photonic crystal tuners. For the

majority of the examples presented, optically induced heating served as a substitute for

resistive heaters. In the case of particle trapping, optothermal effects provide an

alternative to static physical barriers. Optically induced heating was selected for common

reasons. Power delivery through optical transmission is a contactless method that

eliminates the need to integrate an external power source, which can be often be several

times larger than the device being powered. Optical heating is also dynamic, enabling

reconfigurable delivery schemes. Finally, an optical heat source can simplify fabrication

while maintaining precision, a feature that becomes important during further

minimization.

11

CHAPTER 3

CELL TRAPPING AND CULTURE

3.1 Introduction

One of the remaining challenges for microfluidic culture systems is the ability to

culture specific, individual cells. The isolation and culture of specific, individual, and

potentially rare cells of interest, such as cancerous cells, can be used to test the

effectiveness of drugs, which can lead to the development of treatments that are better

suited for their target. The existing techniques for seeding 3D cultures mentioned in

Section 1.1 all involve the use of a bulk cell suspension. The selection of cells can be

performed as a separate step, provided that there is some distinguishable trait to sort by.

Passive sorters rely on deviations in the particle flow streams caused by non-contact

means (e.g., hydrodynamic, dielectrophoretic, or optical forces), whereas active sorters

use a physical mechanism to separate the particles [47]. Methods of actuating

microfluidic cell sorters are diverse, and include the use of hydrodynamic flow [48],

magnetic [49], dielectrophoretic (DEP) [50], [51], acoustic [52], and optical forces [53],

[54], among others. Some of these mechanisms such as DEP [55] and optical trapping

[56], in addition to others including physical micromanipulators [57], [58], can also be

applied to the manipulation of single cells.

By employing one of these methods, it is possible to pre-screen the cells prior to

culture; however, the question of how to recover and seed the cells arises. Cells isolated

using contact-based manipulators can experience problems releasing due to surface

forces, but can be preferable to non-contact manipulators because they can exert stronger

forces necessary to move larger particles [58]. When isolating specific cells for culture, a

preferable method would be to streamline the process by simultaneously selecting the

cells of interest and seeding them into a scaffold. Minimizing the manipulation required

to arrive at the culturing stage also simplifies the equipment necessary to carry out the

process.

The development of a method that integrates both cell selection and seeding

processes into a single step is the focus of this chapter. With the aid of a microscope to

visualize the cell sample, target cells can be identified and captured to create 3D cultures.

12

In the following sections, a microfluidic system is presented where specific cells can be

isolated from suspension through immobilization in a 3D scaffold.

3.1.1 Background

Hydrogels constitute one class of materials used in the preparation of 3D culture

scaffolds. A hydrogel consists of a network of homopolymer or copolymer chains that are

capable of swelling with water or other aqueous solutions. The 3D structure of a hydrogel

is maintained due to crosslinking between the polymer chains. The hydrogel structures

can be formed by chemical crosslinks including covalent and ionic bonds, as well as

physical crosslinks arising from hydrogen bonds, van der Waals forces, or the

entanglement of the polymer chains [59], [60]. Covalent and ionic bonds are stronger due

to intramolecular interactions, whereas hydrogen bonds and van der Waals force are

based on weaker, intermolecular interactions.

Some hydrogels exhibit reactions to environmental stimuli such as temperature,

pH, electrical current, or light [60]. Depending on the strength of the crosslinks, the

reaction may cause a hydrogel to undergo a sol-gel phase transition or a volume phase

transition consisting of shrinking and swelling.

Poly(N-isopropylacrylamide) (PNIPAAm) is a type of thermoresponsive hydrogel

that has gained much attention for its usefulness in biomedical applications. PNIPAAm is

biocompatible, but usually not biodegradable [61]. At temperatures below its lower

critical solution temperature (LCST), approximately 32 ˚C, PNIPAAm polymer chains

are soluble, forming solutions that are transparent in appearance. Raising the temperature

of the solution above LCST rapidly induces a reversible sol-gel transition, resulting in the

formation of an opaque hydrogel. At the microscale, heating causes dehydration of the

hydrophilic polymer chains and their subsequent collapse [62] [63]. The collapsed chains

form hydrophobic globules that immediately aggregate.

Permanent PNIPAAm hydrogels can be prepared by crosslinking with a

photoinitiator. The polymerized PNIPAAm will still undergo a transition at LCST, but in

this case the reaction is constrained to swelling and shrinking of the entire hydrogel [64]-

[66]. However, as with other photoinitated polymerizations [31], [32], [67]-[69], the

crosslinking is usually irreversible, exposes the cells to a dose of UV light, and involves

13

the use of toxic prepolymer and photoinitiator agents.

PNIPAAm is of great interest for biomedical applications due to its

biocompatibility, LCST near physiological temperatures, and the sharpness of its

transition between solution and gel [62]. Biomedical applications involving PNIPAAm or

PNIPAAm copolymer hydrogels primarily use its permanently crosslinked forms.

PNIPAAm has been used in cell culture, most notably as a thermoresponsive coating for

2D culture dishes, providing a less destructive way of releasing affixed cells [62].

PNIPAAm and PNIPAAm-based hydrogels are also the focus of research on injectable

scaffolds fostering cell growth [61], [65]. Furthermore, PNIPAAm copolymers have been

studied as a medium for controlled drug release, including an implementation as an

ocular injection tested in rodent models [70], [71]. These applications have demonstrated

the biocompatibility of PNIPAAm from the cellular to the organism level.

In other cell-related work, PNIPAAm was demonstrated as a tool for cell

immobilization [72]. In a two-step process (Fig. 3.1), a single yeast cell was isolated with

optical tweezers from a bulk sample mixed into 10% (w/t) PNIPAAm solution (Fig. 3.1a,

b) and then was immobilized in gel with heat from an indium tin oxide (ITO) thin-film

heater (Fig. 3.1c). After the target cell was effectively gelled in place, the extraneous cells

in the surrounding PNIPAAm solution could be flushed from the chamber and replaced

with growth medium (Fig. 3.1d).

Fig. 3.1. Illustration of cell isolation procedure proposed by Arai et al. [26] (a-b) Optical tweezers are used to maneuver a single cell toward the location of the ITO heater (yellow strip). (c) ITO heater is turned on, causing localized gelation and trapping the cell. (d) Extraneous cells not caught in the gel are flushed away.

The 3D culture system described here adopts a similar approach to that of Arai et

al. [26], but the cell manipulation step is eliminated, and instead of an embedded micro-

heater, computer-generated optical patterns are utilized to induce the sol-gel transition.

Optically induced heating offers more versatility in controlling gelation, since localized

heat is dynamically produced wherever the light is focused and not confined to a fixed

14

location like with a heater. To reduce the light intensity required to raise the solution

temperature to 32 ˚C, an absorbent substrate is used to facilitate the conversion of light to

heat. Since the sol-gel transition is reversible, this mitigates the issue of biodegradability

since the cultured cells can be easily harvested by dropping the temperature below the

LCST.

3.2 Experimental Setup

The optically controlled 3D cell culture system utilizes a microfluidic device with

an optically absorbing substrate that converts the optical patterns into heat. This heat is

used to trigger the sol-gel transition of PNIPAAm solution containing cells, allowing

selective trapping of target cells in a 3D matrix for culturing.

In the experiments presented here, the cell culture device is positioned on the

stage of an upright microscope, on top of an indium tin oxide (ITO) thin-film heater (Fig.

3.2). The ITO heater is optional, and is used to warm the device above room temperature,

but below the LCST, to facilitate the optically induced gelation. The heater is also used

later in the trapping procedure to maintain the hydrogel regions while transitioning away

from optical heating. Fluid is introduced into the chamber using a syringe pump. A CCD

camera mounted on the microscope provides real-time visualization of the procedure. The

thermoresponsive polymer solution containing the cell sample is introduced into the

device.

Optical patterns are directed through a hole in the microscope stage to the device

substrate. The light patterns are created by focusing the output of a commercial projector

(Dell 2400MP) through a 10X objective lens, resulting in an average intensity of 4.07

W/cm2. An alternate light source that can be used to cause gelation is a 635-nm, 10-mW

diode laser. Irradiating the absorbing substrate results in localized heating within the

device.

15

Fig. 3.2. Experimental setup. The microfluidic culture device containing cells suspended in PNIPAAm solution is positioned on a microscope stage. Light patterns from a projector are focused on the bottom of the device, passing through an opening in the stage. Optionally, an ITO thin-film heater can be inserted between the culture device and the stage to facilitate heating.

3.2.1 Culture Device

The culture device consists of a 50-µm-high PDMS microfluidic chamber

reversibly bonded to an optically absorbent substrate (Fig. 3.3). An inlet port and an

outlet port are located at opposite ends of the chamber, and flushing is controlled by

negative pressure-driven flow using an external syringe pump.

The optically generated heat in the culture device is sufficient to increase the

temperature of the PNIPAAm solution above its LCST, thus inducing localized sol-gel

transitions. The dimensions of the hydrogel are determined by the heat distribution

generated by the optical patterns, and the height of the microfluidic chamber. The height

of the microfluidic chamber in these experiments was maintained at 50 µm to ensure

sufficient transport of nutrients and waste products into and out of the gels once cells are

in culture.

16

Fig. 3.3. (a) Schematic of the 3D cell culture device. A microfluidic chamber is imprinted on polydimethylsiloxane (PDMS) using soft lithography, and reversibly bonded to an optically absorbent substrate. Tubing is connected at the inlet and outlet ports. Fluid is introduced at the inlet and is drawn through the device by applying negative pressure at the outlet. (b) A prototype device is shown with fluidic tubing connected at the ports. In this design, the hexagonal chamber is approximately 9 mm x 5 mm; the PDMS piece and accompanying substrate are approximately 30 mm x 15 mm.

3.2.2 Culture Protocol

A typical culturing procedure starts by mixing cells at a desired density with a

PNIPAAm solution (Fig. 3.4). The suspension is introduced into the microfluidic

chamber. Optical patterns are then used the form hydrogels around the cells of interest.

After trapping the desired cells, the remaining sample is flushed by flowing buffer

solution through the device while keeping the gelled region above the LCST. After this

point, the optical patterns can be turned off, as long as the device itself is kept above the

LCST. The ITO heater on the microscope can be used to maintain the temperature above

the LCST, but the device should eventually be transferred to a cell incubator, which

provides a more stable temperature, and better environment for long-term cell growth.

The cells can then be cultured. The microfluidic chamber is designed to be removable,

17

providing direct access to the hydrogel, and making subsequent culture steps more

similar to that of a standard 3D culture.

Fig. 3.4. Optically controlled 3D culturing process. (a) The cell sample in hydrogel solution is introduced to the device chamber. (b) Desired cells are trapped in PNIPAAm hydrogel (cylindrical volume) using localized optical heating. (c) Undesired cells that remain in solution are flushed from the device. (d) While maintaining the hydrogel above the LCST, the PDMS chamber is removed, and the device substrate with the patterned 3D hydrogel is placed into culture media to allow cell growth.

3.3 Materials and Methods

3.3.1 Thermoresponsive Polymer

A thermoresponsive solution comprised of 10% (w/v) PNIPAAm (MW 19,000-

26,000, Sigma-Aldrich) dissolved in deionized (DI) water or phosphate-buffered saline

(PBS) (ATCC) was used in the experiments discussed in this paper.

3.3.2 Cells

Madin Darby Canine Kidney (MDCK II) epithelial cells and HeLa cells were

18

cultivated in a standard 2D culture using Dulbecco's Modified Eagle Medium (DMEM)

(Gibco) with 10% fetal bovine serum (FBS) and 1% penicillin streptomycin. The cells

were incubated at 37 ˚C in a humidified 5% CO2 environment. The cells were harvested

using 0.25% Trypsin/0.53mM EDTA, and mixed into a solution of 10% (w/v) PNIPAAm

in PBS for the experiments described here. Viability tests were performed using a

LIVE/DEAD assay kit (Invitrogen) and fluorescent imaging.

3.4 Results

3.4.1 Simulation

The temperature profile in the culture device was simulated using COMSOL

Multiphysics (Fig. 3.5). The heat source in this simulation is a 100 µm × 100 µm square,

representing an optical pattern. The heat source is situated at the top surface of the lower

glass layer, which corresponds to the location of the optically absorbent layer. This

conversion of light to heat is an important parameter in determining the efficacy of the

substrate and ultimately the resolution of the gel that can be achieved. Thus, various

light-absorbent substrates were developed and characterized. The materials used to

fabricate the substrates were chosen on the basis that they were readily available as well

as biocompatible. The primary substrate used consists of a 1.1-mm-thick glass slide

coated with a 200-nm-thick layer of ITO and a 1-µm-thick layer of amorphous silicon (a-

Si). The a-Si layer absorbs near infrared, visible, and UV light, converting it to heat. A

simpler substrate was also tested, consisting of a 150-µm-thick glass coverslip backed

with a layer of black vinyl tape (3M Corp.). Additionally, absorbing polymeric substrates

were created from PDMS saturated with dark microparticles such as black mineral

pigment and carbon black.

19

Fig. 3.5. Simulated temperature (degrees Kelvin) in the culture device. In this model, two glass layers sandwich a 100-µm thick layer of water. The heat flux is situated at the bottom of the water layer in a 100 µm x 100 µm square area. 3.4.2 Substrate Characterization and Selection

In order to precisely select single, individual cells for culture, the hydrogel

resolution needs to be optimized. While it is desirable to obtain the smallest hydrogel

area necessary to fully encompass a cell, the area of hydrogel should also remain

relatively constant over the short period of time between trapping and flushing steps. If

the heated region of the substrate continues to spread outward from the point where the

light pattern is aimed, the expanding gelation may potential trap additional, undesired

cells in the immediate area.

For this reason, the substrates were compared by measuring the area of the

hydrogel while exposing the device to the smallest circular light pattern capable of

producing gel within the first 30 seconds of exposure. First, the minimum size of the

optical pattern required to induce gelation was determined for each substrate. The time

point at which hydrogel initially appeared varied across the substrates. The growth of the

hydrogel was measured over a total exposure time of 60 seconds to include the effects of

continued hydrogel formation with time (Fig. 3.6). The diameters of the circular optical

patterns used for the different substrates were: mineral pigment (450 µm); electrical tape

20

on a glass coverslip (550 µm); carbon black (420 µm); a-Si (580 µm).

Although the a-Si substrate did not produce the best size stability over time, it was

selected for use in the subsequent cell-trapping experiments presented here, as it did not

interfere with the quality of the microscope images of the microfluidic chamber. The

absorptive nature of the other substrates relies on the presence of a dense amount of dark

pigments, which had the effect of hindering the reflective microscopy used. The ability to

visualize the cells is crucial in differentiating anomalous cells from the bulk of a

suspension, so the superior imaging qualities of the a-Si substrate made it easier to work

with than the other substrates.

Fig. 3.6. Hydrogel area versus time for various optically absorbent substrates. The smallest optically pattern that effectively produced gel was determined for each substrate, using the computer projector light source. The graph shows the average hydrogel area measured across five trials and the standard deviation for each. A two minute delay between subsequent trials allowed sufficient time for the device to return to room temperature.

The resolution of this technique can be improved by replacing the computer

projector with a laser operating in the near-infrared, visible, or UV wavelength range.

Hydrogel formation using the computer projector light source is limited by the intensity

of the light from the lamp in the projector. The relatively low intensity requires larger

0

0.02

0.04

0.06

0.08

0.1

0.12

0.14

0 10 20 30 40 50 60

Gel

Are

a [m

m2 ]

Exposure Time [s]

Carbon black

Pigment

Tape

a-Si

21

illuminated areas to create sufficient heat to trigger the PNIPAAm gelation. Thus, the

resolution of hydrogel formation was examined further using the a-Si substrate and a

635-nm, 10-mW diode laser as the optical source. The laser was focused down to a spot

of approximately 108 µm2. The substrate was irradiated for 30 s at various laser powers,

and the resulting hydrogel area was measured (Fig. 3.7). The smallest resultant gel area

was 182 µm2. To remove the effect of accumulated substrate heating, measurement at

each power level was performed at a different location in the microfluidic chamber. The

gels formed with the laser have a smaller area than those produced using the computer

projector. However, since the laser is focused into a single point, it does not provide as

much flexibility as the projection of computer-drawn patterns. Thus, the trapping

experiments utilized the computer projector setup, although more flexibility would be

possible by using holographic imaging or spatial light modulators, similar to those used

in holographic optical tweezers [73].

Fig. 3.7. Hydrogel area as a function of power, using a 635-nm diode laser as the optical source. The area of the laser spot was approximately 108 µm2. The substrate was irradiated a period of 30 seconds to generate each data point. The graph shows the average hydrogel area measured across five trials and the standard deviation of the measurements. Minimum gel resolution achieved was 400 µm2. 3.4.3 Mammalian Cells

The trapping capability of the culture system was tested on two types of

mammalian cells. Single and multiple-cell trapping as well as parallel trapping were

0

0.01

0.02

0.03

0.04

0.05

0.06

0 2 4 6 8 10 12 14

Gel

Are

a [m

m2 ]

Power [mW]

22

demonstrated with MDCK II cells. The cells were added to PNIPAAm/PBS solution,

resulting in a cell density of approximately 1×105 cells/ml and 10% (w/v) concentration

of PNIPAAm. The MDCK II cells were selectively trapped using optically induced

hydrogel formation (Fig. 3.8). Initially, the cell solution was introduced into the

microfluidic chamber, and target cells were selected (Fig. 3.8a, d). Under no-flow

conditions, appropriate light patterns were activated to trigger hydrogel formation in the

immediate surrounding area, trapping specific cells. Single cells can be trapped, (Fig.

3.8a-c) or multiple cells (Fig. 3.8d-f). Cells can be trapped in parallel by using multiple

light patterns (Fig. 3.8e). Flowing PNIPAAm/PBS or buffer into the device flushes cells

that are not trapped, leaving behind only the cells within the optically patterned gels (Fig.

3.8c, f). At this point, the light patterns can be turned off as long as the device

temperature is maintained above 32 ˚C.

23

Fig. 3.8. Gelation of hydrogels using optical control, demonstrating the trapping of MDCK II cells in 3D hydrogels. (a) Initial cell distribution. The two target cells are circled in the inset. Other cells will be flushed away. The inset scale bar is 100 µm. (b) A light pattern (bright circle) is used to trigger hydrogel formation (dark circular area), trapping the target cells. Cells not trapped in the gel are flushed away. (c) The two target cells remain in the 3D hydrogel. (d) Initial cell distribution when using parallel optical patterns to trap MDCK II cells in 3D hydrogels. Four target cells are circled in the inset. Other cells will be flushed away, except those trapped by the second optical pattern. The inset scale bar is 100 µm. (e) Light patterns (bright circles) are used to trigger hydrogel formation (dark circular areas), trapping the target cells while free cells are flushed away. (f) The four target cells remain in the 3D hydrogel.

In a separate experiment, the viability of MDCK II and HeLa cells immediately

after a trapping procedure was assessed using a fluorescent LIVE/DEAD assay. For this

procedure, the cells were suspended in a 10% (w/v) PNIPAAm in PBS solution. The cells

were irradiated with a circular light pattern 200 µm in diameter for approximately 20

minutes while trapping and flushing took place (Fig. 3.9a, 3.10a). Pure PBS was used to

24

remove the cells not trapped in the gel from the device, and then PBS solution containing

the viability assay components was flushed in. As soon as flow subsided, the light pattern

was removed while the temperature of the ITO heater was raised above the LCST to

compensate for the loss of heat. The cells were incubated for 25-30 minutes as the stain

was allowed to permeate the hydrogel.

Cell viability of a MDCK II sample was approximately 100%, with no dead cells

visibly fluorescing (Fig. 3.9b). In the experiment with HeLa cells, a little over 80% of the

cells entrapped in the hydrogel fluoresced green, signifying that the cells survived the

process (Fig. 3.10b). Some cells at the border of the hydrogel exhibited the red

fluorescence indicative of cell death (Fig. 3.10c). Overall, it was expected that the

trapping procedure should have an insignificant effect on viability, as PNIPAAm is

biocompatible [61], [62], [65], [66], [70]-[72]. Furthermore, the cells are exposed to an

optical intensity of only 0.17 W/cm2, as most of the incident light is absorbed by the

substrate. This intensity should not directly cause increased DNA fragmentation [74].

These properties make our technique safe for the culturing of specific cells in 3D.

Fig. 3.9. Cell viability testing using LIVE/DEAD assay on trapped MDCK II cells. (a) Initial distribution of MDCK II cells after 50 minutes in an optically patterned hydrogel. The starting cell density was approximately 1×105 cells/ml in solution. The dashed circle encloses the area that was gelled. The circular objects that are visible are the cells. (b) Fluorescent image of the same cells from (a), demonstrating approximately 100% viability. Cells lying outside the gel were not taken into account. Live cells fluoresce green, while dead cells fluoresce red. Fluorescent cells appear as bright spots in these images.

25

Fig. 3.10. Cell viability testing using LIVE/DEAD assay on trapped HeLa cells. (a) Initial distribution of HeLa cells after 45 minutes in an optically patterned hydrogel. The starting cell density was approximately 1×106 cells/ml in solution. (b-c) Fluorescent images of the same cells from (a), demonstrating approximately 81% viability. Live cells fluoresce green (b), while dead cells fluoresce red (c). Fluorescent cells appear as bright spots in these images 3.5 Discussion

3.5.1 Evaluation of Cell Trapping on Viability

Several factors can affect cell viability throughout the culture process. In the

experiments presented here, viability was only assessed immediately after the trapping

procedure, as long-term culture was not performed. A discussion of possible reasons for

the less than ideal (100%) viability is included here. It should be noted that the results are

not statistically significant; to get a better idea of actual performance, viability tests

should be conducted across multiple standardized trials.

First, environmental factors can have an adverse affect on cell viability. During

the loading stage, the cell suspension is kept at temperatures lower than 32 ˚C to prevent

the solution from gelling. Since temperature is used to trigger the hydrogel formation and

maintain it throughout the trapping procedure, it could also influence cell viability. The

heat produced by the projector alone should not be high enough to negatively affect the

cells. The transition between switching over the control of the heat from the projector to

the ITO heater was less regulated and could possibly introduce temperature fluctuations

that are not healthy for the cells.

Second, although less understood, is the interaction of the cells with the

PNIPAAm scaffold. Cells are present within the polymer network while it undergoes the

collapse and dehydration associated with the sol-gel transition. It seems possible that

26

compression forces could arise during the reorganization of the polymer structure, but

prior literature on this was not found. The release of water from the scaffold could also

introduce shear stress on the cells from the fluid flow.

An isolated study investigated the cytotoxicity of non-crosslinked PNIPAAm

[63]. PNIPAAm (MW = 156,000) was synthesized from NIPAM monomer and dissolved

in Hank’s balanced salt solution (HBSS). Caco-2 cells incubated in 0.5% (w/v)

PNIPAAm solution for three hours at 37 ˚C showed decreased viability in comparison to

those maintained at 23 ˚C. Viability dropped to levels under 80% after incubation for 12

hours. Although this implies that PNIPAAm heated above LCST and the length of

incubation time have an effect on cell viability, there are too many variables to establish a

direct correlation to the results presented here (e.g., the molecular weight was greater by

a factor of six and a much lower concentration of PNIPAAm solution was used). Further

study should be done on the PNIPAAm solution used in the culture device. Switching to

a different molecular weight would probably not be an issue; however, the concentration

of PNIPAAm should remain relatively high since trapping and flushing is reliant on the

difference in viscosity between regions of gel and solution.

3.5.2 Considerations for Long-term Culture

For extended culture periods, the permeability of the scaffold to nutrients and

waste metabolites will determine the sustainability of the culture and is an active area of

research. The optimum permeability of tissue culture scaffolds is dependent on a several,

sometimes competing, factors. Requirements can vary based on the type of cell, due to

differences in metabolic requirements. Permeability is also application dependent, such as

in the case of transplantation, where the matrix isolates the cells from the host immune

response, or bioreactors, where it is desirable to harvest cell-derived products.

Another consideration is the variability of extracellular matrix requirements of

different types of cells. Engler et al. demonstrated that the optimal substrates to support

the normal function of cardiomyocytes (cells found in the myocardium of the heart)

optimally have an elasticity that closely resembles the natural ECM [75]. In vivo,

cardiomyocytes adhere to a collagen-based extracellular matrix, providing some degree

of compliance. Cardiomyocytes grown on non-compliant substrates such as glass and

27

lipid-coated polystyrene ultimately lose their normal characteristics including the ability

to contract rhythmically. The normal behavior of the cardiomyocytes could be sustained

for longer periods on more compliant substrates. The authors acknowledge that other

intercellular interactions could change the actual strain experienced by the cells; however,

it was suspected that less compliant matrices resulted in overstretching of the cells,

rendering contractions less efficient.

3.6 Conclusion

In this work, a cell culture system was demonstrated with the ability to trap cells

of interest in reversible 3D scaffolds comprised of a thermoresponsive hydrogel,

PNIPAAm. The culture system utilizes a simple microfluidic device that acts as a

repository for the hydrogel solution containing the cell sample while the trapping

procedure is performed. Optically generated heating is used to induce localized gelation

of cells within the device, enabling the trapping of specific cells. Two possible methods

of generating localized heating at the cellular level were shown: using a modified

commercial projector as a source of light, and using a diode laser. The projector offers a

higher degree of flexibility for pattern generation, whereas the higher intensity of the

laser provides a stronger heat source, and can thus create gels with smaller areas. The

intensity of either source is not at a level where it is expected to damage the irradiated

cells. To test the trapping capabilities of the culture system, MDCK II and HeLa cells

were used. With MDCK II cells, single and multiple cell trapping was achieved, using a

single light pattern or multiple simultaneous patterns. Preliminary viability tests show

that over 80% cells survive the trapping procedure. Future work includes performing

further viability tests at later time points during the cell cultivation process.

28

CHAPTER 4

BUBBLE MICROROBOTICS

4.1 Introduction

Microrobots capable of micromanipulation and microassembly have vast potential

in furthering biomedical technology. One area of biomedical research that has gained

much interest recently is the 3D culture of artificial tissues and organs in vitro. In 3D

culture, cells are grown on a scaffold that mimics the natural extracellular matrix.

Through the use of microfabrication techniques, small blocks of cell-laden scaffold called

microgels can be created [76]-[78]. Acting as functional blocks of tissue, the microgels

can be assembled into desired shapes or patterns, mimicking how groups of tissues are

organized during tissue formation in vivo [77]-[79]. Of the several techniques used to

assemble these microgels [77]-[83], only physical manipulation is capable of assembling

specific individual microgels [80]. The use of wireless microrobots to physically

manipulate and assemble microgels is a viable option.

4.1.1 Background

Limitations in the ability to scale down electrical and mechanical components

used in conventional robotics make the locomotion of wireless microrobots a challenge.

Current microrobotic systems utilize a diverse assortment of control and power delivery

mechanisms, including electromagnetic [84]-[92], electrostatic [93], and optical actuation

[94], all of which utilize resourceful methods to overcome the scaling constraint.

Locomotion at the microscale also takes on a different flavor, with modes of translation

ranging from simple dragging to stick-slip or biomimetic motions. However, for a

micromanipulation technique to be useful in many applications requiring microassembly,

it is necessary to scale the system to process large numbers of objects. One way to meet

these demands would be to develop a multi-agent microrobotic system, enabling the

operation of multiple microrobots in parallel.

A multi-agent system refers to a system that hosts two or more interacting,

intelligent components [95]. Macroscale multi-agent robotic systems are able to

incorporate some type of hardware-based onboard intelligence, allowing differentiation

29

and individual addressing among a batch of otherwise uniform robots. Unlike their

macroscale counterparts, microrobot designs afford very little onboard space, if any.

Because of the size constraint, intelligence—in addition to power delivery—has to be

provided by some other means.

Currently, only a few of the existing microrobot control systems are able to

support the planar actuation of multiple microrobots in a common operating environment

[91]-[93]. These systems approach the task of enabling multi-agent functionality by

giving the microrobots a “mechanical intelligence” [91]. This means that instead of

addressing the microrobots on a 1:1 basis, heterogeneous microrobots were designed that

can selectively respond to a global control signal.

Frutiger et al. demonstrated the independent actuation of two frequency-selective

microrobots that have mechanical responses at different resonant frequencies [91]. Time-

division multiplexing of the control signal enabled simultaneous, independent locomotion

of two microrobots. Unique responses to different frequencies also allow the stick-slip

microrobots of Floyd et al. to be selectively addressed [92]. Without relying on a

specialized substrate, the microrobot designs could be selected based on differences in

both geometry and magnetization. Among a group of three microrobots, one could be

independently addressed while the others moved in parallel, coupled together in groups

of two or three. Structural variations in the steering arm of the scratch-drive actuated

microrobots demonstrated by Donald et al. allowed individual addressing [93]. Assembly

of up to four of these microrobots was demonstrated; however, further scaling is limited

due to prohibitive voltage separation requirements and breakdown of the electrode

insulation in the substrate.

In general, the techniques for decoupling microrobot responses to a single control

signal require precise control systems and the tuning of individual microrobots. This can

be challenging when expanding the number of microrobots in a system. In contrast, the

bubble microrobot system discussed in this chapter uses a completely different approach

to controlling multiple microrobots, described in detail in section 3.4. Here, optical

patterns are used to actuate sub-millimeter air bubbles in a fluidic environment. Localized

temperature gradients resulting from the conversion of the optical patterns to heat drive

the motion of the bubble microrobots. Optical control is attractive due to the versatility of

30

the “control signals” (light patterns) that can be produced and the relative ease of

implementing this type of control system. Moreover, the system enables independent

addressing of microrobots, actuation in series or parallel, and coupled or batch

addressing. This microrobotics project was started as an entry to compete in the National

Institute of Science and Technology (NIST) Mobile Microrobotics Challenge [96], [97] a

competition held at the annual IEEE International Conference on Robotics and

Automation (ICRA), but has since expanded in scope to address biomedical applications.

4.2 Experimental Setup and Control System

The bubble microrobot control system utilizes computer-generated optical

patterns projected onto the absorbent substrate of a microfluidic chamber to establish

local temperature gradients (Fig. 4.1). Real-time visualization of the microrobots is

provided via a microscope-mounted CCD camera (PixeLINK PL-B776U and Prosilica

GE680C, Allied Vision Technologies) connected to a computer. The control system is

broken down into three components and described in further detail.

Fig. 4.1. Experimental setup. A fluidic chamber containing the microrobot(s) and other micro-objects rests on a microscope stage. An opening in the stage allows focused light patterns from the projector and optics beneath the stage to illuminate the device substrate.

31

4.2.1 Optical Patterns

Light emitted from a modified computer projector (Dell 2400MP and ViewSonic

PJD6241) serves as the source of the optical patterns. Different configurations are

designed and animated in software, providing a versatile and affordable means for

defining highly programmable control signals. Minimizing optics located beneath the

microscope stage are used to focus the projector output on the absorbing substrate. The

optical patterns are converted to heat by absorption in the substrate. The resulting thermal

gradient is used to actuate the microrobot.

4.2.2 Microrobot Chamber

The bubble microrobots are actuated in an enclosed microfluidic chamber built on

top of the substrate. The optically absorbing substrate consists of a 1-µm-thick layer of

amorphous silicon (a-Si) coated over a 200-nm-thick layer of indium tin oxide (ITO) on a

1.1-mm-thick glass slide. Using an absorbing substrate reduces the light intensity

required to cause an increase in the localized temperature. Empirical measurements show

that a rise in temperature of approximately 1 ˚C can be achieved by with the projector as

a light source. It should be noted that the substrate has not been optimized for microrobot

performance, as it is originally intended for optoelectronic tweezers (OET). It is possible

that a similarly absorbent substrate could be substituted for the a-Si based one; however,

alternative substrates should be uniformly absorbent, smooth, free of artifacts that affect

visibility or be confused with micro-objects, and keep biological compatibility in mind.

The microfluidic chamber is composed of a glass slide bonded to the absorbent

substrate. Spacers between the substrate and the glass slide define the height of the

chamber. The assembled device is leveled on the microscope stage, as the bubbles will

drift if any tilt is present.

4.2.3 Control Software

Various graphics were generated and controlled using programs written in

Processing, an open source programming language and integrated development

environment [98]. Running a program in the Processing development environment opens

a window that displays the graphics specified within the code. Graphics defined within

32

the Processing code can range from simple to complex; they can also be animated or be

responsive to input from peripherals such as a mouse or keyboard. An image of the

display window is sent to the projector, converting the graphics into optical patterns used

to drive the microrobotic system.

The control programs used in the experiments presented in this paper create white

circles of various diameters against a black background, maximizing contrast. Projecting

one or more spots onto the substrate generates radial temperature gradients at the

locations where the light is absorbed. The programs contain both predefined animations

of circles traversing the display window as well as circles that can be moved by user

input. The current control system does not support feedback from the microscope, thus

some user control is necessary. Currently, the control software relies on mouse and

keyboard input to handle different events including spot generation, deletion, and

movement.

4.3 Actuation Mechanisms and Theory

Optically induced thermocapillary effects drive bubble microrobot locomotion.

This method of actuation has been demonstrated on gas bubbles in oil [99], [100]. When

light patterns are focused onto an optically absorbent substrate, the light is converted to

heat, establishing radial temperature gradients in the fluid. If a bubble is present, these

temperature gradients generate surface tension gradients along the bubble interface and

thermal Marangoni (thermocapillary) effects that give rise to fluid flow. As a result, a gas

bubble in oil exhibits a net movement toward the region with the highest temperature,

where it is then stably trapped. Vela et al. describes a second phenomenon, natural

convection, as contributing to the flow, in work focusing on convection-based particle

manipulation in water [101]. The natural convection alone produced forces similar to

optical tweezers, but this is a different effect from the one described here. In the work by

Vela et al., fluid at the focal point of a 1480-nm laser reached a temperature near boiling.

In the bubble microrobot system, light absorbed from the projector causes an

approximately 1 ˚C increase. Therefore, the effect of natural convection is neglected, and

the system is modeled on thermocapillary flow.

In previous work, the thermocapillary flow profile at the air-oil interface of a

33

bubble was simulated using finite-element modeling software (COMSOL Multiphysics)

[100]. The chamber was defined having height of 300 µm. A fixed bubble with a 125-µm

radius was used. The light pattern was modeled as a saturated Gaussian heat distribution

with a full-width at half-maximum of 424 µm [102]. The peak temperature of the heat

source was set to 300.8 K, matching the empirical value obtained when measuring light

pattern incident on the substrate using a 50-µm thermocouple element (Omega

Engineering, Inc.). The boundary condition at the interface can be described by two

equations:

𝜂 !"!𝒏

= 𝛾!!"!𝒕

(1)

𝛾! =!"!"

(2)

In equation (1), n and t are the unit vector normal and the tangent to the interface,

respectively, 𝜂 is the dynamic viscosity, u is the tangential component of the fluid

velocity vector at the liquid/air interface, and T is the temperature [103]. Equation (2) is

the derivative of the surface tension with respect to temperature, where 𝛾 is the surface

tension [104].

Illuminating a point on the substrate with an optical pattern produces temperature

gradients in both horizontal and vertical directions. Therefore, the simulation was broken

down into two models describing the respective components of the fluid flow: parallel to

the substrate, referred to as horizontal thermocapillary flow, and perpendicular to the

substrate, or vertical convective flow. In the following simulations, a fixed air bubble is

treated as the frame of reference. The bubble is static and flow-induced deformation is

not considered. Empirical observations have shown that the bubble makes a contact angle

of 180˚ with the substrate. Although the simulation describes the flow profile around a

single bubble, parallel bubble actuation has been verified experimentally.

34

Fig. 4.2. Simulations of the thermocapillary flow around a stationary air bubble. (a) A heat source at one end of the bubble produces a temperature gradient along the air-oil interface, causing horizontal thermocapillary flow. (b) The same heat source in (a) positioned at the center of the bubble induces vertical convective flow, shown in this side view of the fluidic chamber.

Using a heat source with maximum temperature at x = 250 µm, the simulation of

horizontal thermocapillary flow illustrates how temperature and surface tension gradients

drive horizontal fluid flow from the heated side of the bubble to the cooler side (Fig.

4.2a). As a result of the flow, the bubble exhibits a net movement toward the heat source.

The motion of the bubble will cease when it is centered at the maximum of temperature

gradient.

The vertical convective flow is caused by the temperature gradient that extends

upward from the substrate to the surface of the fluidic chamber. This flow is present in

the fluid surrounding the heated bubble, even when the bubble is centered about the heat

source. The vertical convective flow is normal to the substrate near the bubble interface,

but is part of a larger circulatory flow (Fig. 4.2b). Objects that have a higher density than

the liquid phase can be pulled toward the bubble by the convective flow. These objects

will remain stably trapped near the bubble as long as the vertical convective flow force is

less than the gravitation force. Conversely, objects that are less dense than the fluid end

up following the circulatory flow, and can be repelled, as observed in experiments

35

involving hydrogel beads.

4.4 Materials and Methods

4.4.1 Microfluidic Holding Chamber

Two chamber designs were used in the experiments presented in this paper (Fig.

4.3); the first was built to address the requirements of the microrobotics competition,

whereas the second is a more generalized design that was used to demonstrate additional

capabilities of the microrobot.

Fig. 4.3. Chamber designs. (a) Chamber with epoxied coverslip spacers and competition arenas patterned on substrate. (b) Chamber with Kapton tape spacers containing micro-objects.

The first chamber consists of a 1.1-mm-thick glass slide bonded to the a-Si

substrate with 300-µm-thick spacers separating the layers (Fig. 4.3a). Epoxy (Devcon,

Illinois Tool Works Inc.) was used as the bonding agent and a stack of three glass cover

slips (each approximately 80 µm thick) formed the spacers. The chamber is open-ended

as only the two sides where the spacers are located are sealed with epoxy. At the open

ends, carrier fluid is introduced to the chamber by capillary flow. The height of the

chamber also provides sufficient room for insertion of a thin syringe tip and micro-

objects. For the final devices used during the microrobotics competition, arena structures

were patterned onto the substrate before the chamber was assembled.

The more rigid design of the epoxy chamber was desirable to ensure that height

requirements were met and to help with transportation. Although the chamber can be

reused, it eventually accumulates enough dust and other unwanted particles that the

microrobot performance suffers. The microrobot movement is sensitive to particles on the

surface of the substrate. Cleaning the substrate is difficult for the epoxied fluidic chamber

since it cannot be disassembled once the epoxy cures.

A second, simpler chamber design enables better reuse of the substrate across

multiple experiments due to the lack of a permanent bond (Fig. 4.3b). This chamber is

36

similarly built: a-Si-coated glass serves as the substrate; a glass coverslip replaces the

glass microscope slide as the chamber cover; and two sides of the chamber are left open-

ended. Instead of cover slips, the spacers consist of three layers of double-sided Kapton

tape. The tape provides a way to fix the height of the chamber and also creates a seal

between the layers. The tape adhesive provides a useable, but weaker seal; therefore the

device should be reassembled frequently. Reassembly coincides with the need to clean

the substrate of debris and does not take much effort.

Inside the chamber, the bubble microrobots can be operated within defined spaces

of an “arena” consisting of microfabricated barriers on the substrate. Inclusion of these

arenas was due to the requirements of the microrobotics competition, although they could

also be useful in ordinary microassembly tasks. Based on specifications from NIST, two

arena layouts were drawn to scale in AutoCAD and used to create transparency

photomasks.

Fabrication of the arena structures begins with cleaning a-Si substrate with a serial

wash of acetone, methanol, isopropyl alcohol (IPA), and deionized water. After cleaning,

the substrate was spin-coated with SU-8 photoresist (MicroChem Corp.) to a thickness of

50 µm. Next, the SU-8 was sequentially pre-baked, soft-baked and exposed under a mask

aligner according to product specifications. It then underwent a post-exposure bake

before developing for approximately six minutes with MicroChem SU-8 developer. The

developer was removed with IPA coupled with light agitation and the remaining

structures were blow-dried with nitrogen before undergoing a final hard-bake.

4.4.2 Bubble Microrobot

The microrobot used in this system consists of an air bubble in carrier fluid. Two

types of oil, silicone oil (Dow Corning Corp. 200) and Fluorinert™ FC-40 oil (3M

Corp.), have been tested. In a separate work, laser-actuated air bubbles in aqueous

solutions have also been demonstrated [105]. Bubbles are introduced into the chamber by

extruding a small amount of air into the oil from a fine-tip glass syringe (gastight syringe

models 1001, 1002, and 1005, Hamilton Co. and 33 gauge metal hub needle model

91033, Hamilton Co.). In order to place the bubble past the raised walls of structures built

inside the chamber, it was necessary to use a syringe tip that was thin enough to be

37

inserted over the walls. To release the bubbles, tilting the chamber at an angle so that the

open end is elevated and shaking it a bit will cause the bubbles to rise and exit the oil.

4.4.3 Micro-objects

Various sub-millimeter objects of different materials and geometries were used to

test the micromanipulation and microassembly capabilities of the microrobot. These

include 105-150-µm-diameter glass beads (PolySciences, Inc.), polymer-based triangle

blocks, and hydrogel beads and blocks containing yeast cells. All micro-objects except

glass beads were fabricated in-house.

The second task for the 2011 Mobile Microrobotics Challenge involved packing

triangular blocks within an “alley,” or a narrower section of a T-shaped arena. Right

triangle blocks with legs of 360 µm and 185 µm were fabricated from SU-8 in a modified

version of the arena fabrication procedure described above. Here, a clean silicon wafer

was used instead of the a-Si substrate. The SU-8 was spin-coated at an average thickness

of 100 µm, and was developed for approximately ten minutes. To retrieve the triangle

pieces, the wafer was sprayed with IPA, removing the undeveloped SU-8 and dislodging

the triangles.

Hydrogel beads containing yeast were also fabricated and tested. Agarose is a

biocompatible, thermoresponsive hydrogel that is commonly used in cell culture.

Structurally, agarose is a polysaccharide biopolymer that is derived from natural sources.

It is soluble in aqueous solutions upon heating beyond a particular melting temperature

that varies with the type of agarose. Once it is dissolved, the agarose hydrogel undergoes

a reversible sol-gel transition in response to temperature. Hysteresis in the

thermoresponsive behavior prevents agarose gels from reverting to solution form when

exposed to temperatures between gelling and melting points. Physiological temperatures

fall within this range; therefore the gelation will not reverse during manipulation and

subsequent culture.

A common method of encapsulating cells in agarose is by fabricating agarose

beads. First cells are suspended in agarose solution. Droplets of the suspension can be

created a number of ways, such as extrusion or agitation to form an emulsion in an

immiscible fluid such as oil [35]. Once the beads are formed, the agarose is gelled by

38

dropping the temperature below the gel point. The agarose provides a protective buffer

for the cells during manipulation. A solution of 1.5% (w/w) low-melting-point agarose

(Promega) in deionized water was prepared. Powdered agarose mixed with water was

heated in a microwave for one to two minutes until the agarose was completely dissolved.

The solution was reweighed and adjusted for water lost through evaporation during the

microwaving.

For the bottom-up development of artificial tissues in vitro, agarose-based

microgel blocks can be used in place of beads. Preparation of microgels in various

geometries can be accomplished by using a modified version of McGuigan et al.’s

procedure for making collagen gels with a microfabricated PDMS membrane mold [106].

Transparency masks with the desired microgel patterns are designed and used to pattern a

silicon wafer with SU-8 structures. A thin layer of PDMS is spin-coated onto the

patterned wafer and excess is scraped off of the tops of the SU-8 structures. Once the

PDMS is cured, it is removed from the photoresist structures; the result is a thin

"membrane" that serves as a mold for the agarose blocks. Diced and cleaned portions of

the membrane are laid flat on glass slides. The hydrophobic PDMS is corona treated,

rendering it temporarily hydrophilic to allow the agarose to seep into the holes. Molten

agarose with cells is deposited onto the mold, excess is leveled with a razor, and the

remaining contents are gelled.

Bakers’ yeast (Fleishmann’s) was rehydrated and mixed into agarose solution that

had been cooled to approximately 30˚C. The agarose-yeast mixture was then added to

chilled Fluorinert™ FC-40 oil (3M Corp.) and shaken to form a droplet emulsion. To

form beads, the temperature of the solution was quickly dropped below the gelling point

of the agarose (24 to 28˚C) by refrigerating for approximately five minutes. Bead samples

could then be retrieved by pipet.

4.5 Results

4.5.1 Force on a glass bead

The force produced by bubble microrobots on a 60-µm-radius glass bead in

silicon oil was measured (Fig. 4.3). Glass beads are denser than the oil, so they sink to the

bottom of the chamber. Two mechanisms for manipulating the beads were observed,

39

referred to here as pulling and pushing modes. In the pulling mode, the bead becomes

trapped in the vertical convective flow at the base of the bubble, pulling it inwards toward

the bubble. In this mode, the bead does not have to be in direct contact with the bubble

meniscus since the circulatory flow extends outward into the oil surrounding the bubble.

Conversely, in the pushing mode, the bead is continuously in contact with the bubble, and

is moved by the bubble meniscus.

The effect of increasing the radius of the bubble while maintaining the light

pattern at a constant size was examined. The force exerted on the glass bead was

calculated using Stokes’ Law. Friction with the surface is not taken into account, so the

actual force produced on the bead by the bubble microrobot is likely larger than the

values shown here (Fig. 4.4). The pushing and pulling forces increase until they reach a

maximum for bubble radii between 100 and 200 µm. The decrease in force for larger

bubbles can be partly attributed to the decrease in actuation velocity as bubbles of larger

size are used.

Fig. 4.4. Relationship between the average force exerted on a glass bead and the radius of the bubble microrobot. The radius of the bead being manipulated was 60 µm. The average light intensity of the optical patterns was 5.3 W/cm2 and the light pattern radius was 140 µm. 4.5.2 Microassembly of SU-8 Triangles

During the 2011 Mobile Microrobotics Challenge, the bubble microrobot was

able to move two SU-8 triangle blocks from the starting area into the alley, but did not

40

accomplish the goal of tightly packing the them within the allotted time. The

manipulation and tight packing of three of the SU-8 triangles over the course of six

minutes was presented in a separate work [100]. As the SU-8 blocks are denser than

silicone oil, they settle on the bottom of the chamber rather than float. Friction can be a

problem when the interior of the chamber becomes too dirty; the microrobot or other

micro-objects can become stuck, or have their motion impaired. The problem of friction

is likely exacerbated by the coarseness of the triangles’ surfaces resulting from the

fabrication process.

4.5.3 Parallel Actuation and Microassembly

One application of the bubble microrobots introduced in previous work was

microassembly of hydrogel beads using a single bubble [100]. Parallel actuation of three

microrobots was also demonstrated, although the microrobots were not manipulating any

micro-objects. Here, parallel actuation of multiple microrobots is further explored, and

subsequently combined with microassembly techniques to demonstrate cooperative

microassembly.

Fig. 4.5. Six microrobots are manipulated in parallel. Groups of three microrobots assigned to a single set of controls (a) are moved together forming a 2x3 array (b). The groups are returned to their original configuration (c) and moved together again, this time into a 3x2 array (d). Intensity of the light source was approximately 6.89 W/cm2 and 0.414 W/cm2 after passing through the substrate.

41

Multiple microrobots can be actuated in close proximity to one another. Six

microrobots were simultaneously actuated using keyboard controls, with groups of three

assigned to a single set of controls (Fig. 4.5). Although it is possible to independently

operate each of the microrobots, the task of independent and simultaneous manual control

becomes increasingly difficult with many microrobots. In the current system, the ability

to simultaneously address individual microrobots can be achieved one of two ways: by

using predefined moving light patterns or by manual control of the light patterns using

keyboard inputs. The former method was used in the previous demonstration of parallel

actuation [100]. The latter method requires that enough sets of unique keys are available

to control microrobot movement. Furthermore, a single user can only actuate two

independent microrobots at the same time and still be able to use them purposefully.

Fig. 4.6. Manipulation of a sample of agarose beads containing yeast using two independently controlled bubble microrobots. The initial cluster of beads (a) was pushed into a horizontal line (b), then rotated 90˚ (c-d). Moving the bubbles along the beads straightened them into a single-file line (e-f). The small bright dot in the photos is the projection of the mouse cursor onto the device and the tiny black flecks are stray yeast.

Two independently controlled microrobots were cooperatively actuated to arrange

yeast-laden agarose beads (Fig. 4.6). In this experiment a separate person controlled each

microrobot using keyboard inputs. The agarose beads are less dense than the carrier fluid,

allowing them to float. According to the vertical convection flow model, the floating

42

beads are in the region where convective flow is circulating away from the bubble

interface. However, the beads are large enough that they are not continuously circulated

in the convection, and are instead repelled by the flow surrounding the bubble

microrobots.

Initially, the microrobots were used to clear a region of extraneous beads. The

microrobots were positioned on opposite sides of a small group of beads (Fig. 4.6a), and

then moved toward each other to maneuver the beads into a horizontal line two beads

wide (Fig. 4.6b). The microrobots then rotated the line of beads 90˚ counterclockwise

(Fig. 4.6c, 4.6d). The bubbles were then brought closer together and moved along the

line, straightening the beads into a single-file line (Fig. 4.6e, 4.6f).

Fig. 4.7. A 66-µm-radius glass bead is passed between three independently controlled microrobots. The bead is passed from the middle to the top microrobot (a-b), from the top to the bottom (c-d), and from the bottom back to the middle (e-f). Intensity of the light source was approximately 6.89 W/cm2 and 0.414 W/cm2 after passing through the substrate.

Three microrobots were used in parallel to perform a cooperative bead "handoff"

routine (Fig. 4.7). The upper and lower microrobots were independently controlled by

keyboard while the central microrobot followed an automated path traveling from left to

right and back again along a horizontal line. In a combination of pushing and pulling

43

modes, a single glass bead was passed between the bubbles, following a sequence of

middle-top-bottom-middle.

4.6 Discussion

4.6.1 Variations due to operating environment

Whether or not micro-objects will float in a particular carrier fluid determines the

mode(s) of microrobot operation. Selection of the environment may be important

depending on the type of micromanipulation desired (e.g., sorting, microassembly, etc.)

Empirical results show that both pushing and pulling modes can be used to

manipulate objects with a density greater than the carrier fluid. Although the pulling

mode is noticeably weaker, having both modes available can be useful for obstacle

navigation, correction during microassembly tasks, or in cooperative microassembly.

When the objects under manipulation are less dense and float, as in the case of

agarose beads in FC-40, only the pushing mode is observed. Unlike the pushing mode

exhibited in the glass bead and triangle manipulation, the agarose beads remain at a

minimum distance from the bubble at all times. Manipulation of floating micro-objects

appears to be more fluid, due to reduced friction experienced by the micro-objects.

However, floating can be a disadvantage for microassembly, since the objects will not

remain in place without extra supportive structures to tether them.

4.6.2 Considerations for further development

Advantages of the bubble microrobot system include relatively few hardware

requirements, uncomplicated microrobot and platform fabrication, and a straightforward,

graphics-based approach to control software. This system is unique in that the microrobot

itself does not require special material, just air. New microrobots are generated on-

demand and can be replaced indefinitely. Different exercises in micromanipulation have

illustrated the versatility of the bubble microrobots as a scalable, multi-agent system. The

control mechanism utilizes 1:1 addressing—one light spot to control one bubble—which

makes scaling up easier than the other three control systems previously mentioned.

Instead of extensive work on developing and controlling multiple unique designs, the

44

methods described in this paper can possibly meet scaling goals with further work in

characterization, optimization, and automation of the microrobot actuation.

To facilitate expansion to larger scale operation, certain aspects of the

microrobotic system can be improved. Limitations of the bubble microrobots discovered

in practice were primarily related to the behavior of air bubbles in carrier fluid. In the

following paragraphs, potential solutions to these issues are discussed.

One issue is variation in bubble size. Extrusion of air through a fine-tipped

syringe is currently used since it does not require any specialized equipment. The

drawback of this method is inconsistency in bubble size, evident in the parallel actuation

images. Adoption of an alternative method that offers increased precision should be

considered. Kasumi et al. used a 5-µl syringe to prepare bubbles of equal size by

alternating the withdrawal of oil and desired volumes of air into the syringe [107]. In

practice, the peak actuation velocity of a bubble is influenced by the relationship between

the bubble radius and the radius of the light spot controlling it. Peak velocity appears to

occur around when bubble and light spot are of similar size. Accounting for fine

variations in bubble radius among multiple bubbles in which similar actuation velocities

are desired is probably best handled by automation. Incorporating image analysis of the

real-time feed from the camera, an automated routine could be used to select

appropriately sized light patterns based on the bubble radius.

A second issue is that the formation of bubble microrobots can only occur in

carrier fluids that have enough surface tension to maintain bubbles for reasonable

amounts of time. For the current setup, the carrier fluid was limited to oils, which are

rather viscous fluids. Increasing the intensity of the light source as demonstrated with

laser-operated microrobots enables the use of the microrobots in less viscous aqueous

solutions [105].

A third issue that can disrupt operation is the gathering of multiple bubbles

around a single light spot. Thermal gradients induced by the heat source can draw in

bubbles from within the vicinity of the source through horizontal thermocapillary flow.

The circulating thermocapillary flow about the bubbles themselves can be sufficient to

attract other bubbles. The theoretical and experimental study performed by Kasumi et al.

on pairs of air bubbles in silicone oil on a heated wall showed that bubble aggregation is

45

due to thermocapillary effects and that the relative motion of the bubbles scales linearly

with temperature gradient, bubble size, and viscosity-1 of the liquid phase [107].

Increasing the bubble radius alone resulted in higher aggregation velocity, implying that

there is some difference in the thermocapillary flow profiles; however, it cannot be

determined from these results whether this is due to increased flow velocity, how far the

flow extends into the surrounding fluid, or some combination of both. Intuitive collision

prevention that takes into account the relationship between bubble size and flow profile

can help avoid bubble aggregation situations. Unfortunately, the microrobot system is

more complex than the one used by Kasumi et al. since instead of a uniform vertical

temperature gradient, localized temperature gradients with different areas are used. After

the relationship is established, a minimum separation distance between neighboring

bubbles as a function of bubble size can be included in a collision prevention routine. In

the absence of automation or if it were to fail, the aggregated bubbles can be separated

manually. From experience, temporarily removing the heat source results in allowing

touching bubbles to coalesce. In these instances, it is better to use additional light spots to

separate the cluster and isolate the bubbles.

Although the use of bubbles rather than traditional solid structures is beneficial

for several reasons, a fourth issue is that bubbles under manipulation have often been

observed to enlarge over an extended period of time, at times becoming too large to use.

The cause of the size instability is not fully understood, and is currently under

investigation.

A final problem arises when solid structures render bubbles immobile. Rough

walls, substrate defects, and debris are all possible structures that can cause a bubble to

become stuck. This does not occur all of the time and not all obstacles will impede

movement. Stronger convection may allow the bubbles to overcome the friction forces.

4.7 Conclusion

In this chapter, optically actuated bubble microrobots were discussed. Actuation

has been demonstrated in two types of oil and manipulation of glass beads, SU-8 blocks,

and agarose beads and blocks were presented. The ability to actuate multiple microrobots

in parallel is a benefit of this system. Here, independent, simultaneous control of multiple

46

microrobots within close proximity of each other was demonstrated. Preliminary results

showing the effectiveness of parallel actuation for microassembly included two

microrobots working in unison to assemble hydrogel beads and three microrobots passing

a glass bead around.

Future areas for improvement of the bubble microrobot system include improving

management of bubble size, development of autonomous control using image feedback

integrated with the control software, and characterization of the dynamics of the multiple-

bubble system. Related work pursued in parallel and published separately involves the

manipulation of micro-objects in aqueous fluidic environments using laser-driven

bubbles, thus improving compatibility with biomedical applications [105].

47

CHAPTER 5

CONCLUSION

In this paper, two systems driven by optothermal effects with applications in

routine cell-handling processes were presented. Operation of the similarly designed

microfluidic devices relies on the same control system and substrate for optical-to-

thermal conversion. The common control system comprises a customized commercial

projector, minimizing optics to direct and compress the light, and software that creates

the images to be projected, and is built around a microscope with a mounted CCD

camera. A substrate consisting of a-Si and ITO coated glass that was originally developed

for OET was used.

Chapter 2 introduced several MEMS technologies that also utilize optothermal

effects. A variety of applications exist; of these, sensors, actuators, manipulators, and a

tunable photonic crystal device were discussed. The majority use optically induced

heating as a replacement for integrated resistive heaters. In all but the sensor example,

optothermal control is favored because it is a dynamic, non-contact method of power

delivery, usually simplifies implementation, and is highly reconfigurable, especially in

comparison to embedded heaters.

In chapter 3, a system for cell trapping and subsequent culture was designed and

tested with MDCK II and HeLa cells. Contained within a simple microfluidic chamber, a

solution consisting of PNIPAAm dissolved in aqueous solutions responds to localized

heating created upon irradiation of the substrate with optical patterns, undergoing a sol-

gel phase change. Gelation occurs at approximately 32˚ C, a little lower than normal body

temperature. If cells are suspended in the PNIPAAm solution, directing the light patterns

at regions around cells of interest selectively traps these cells in gel. The PNIPAAm

hydrogel doubles as a scaffold for 3D cell culture while the remainder of the PNIPAAm

solution is replaced with buffer, and ultimately media to support long-term cultivation.

In chapter 4, the focus shifts to an optothermal system for controlling bubble

microrobots in oil that are capable of micromanipulation and microassembly tasks.

Although applicable in a broader range than the first device, the bubble microrobot

system can be used to assemble cell-laden microgels as units of tissue culture. Radial

48

temperature gradients established by irradiation of the substrate with animated optical

patterns induce thermocapillary flows that drive microrobots locomotion. The

microrobots can be used to push micro-objects, and sometimes pull them, depending on

whether the objects float or not. The microrobots were tested on four types of micro-

objects: glass beads, easy to transport because they roll; SU-8 blocks, which proved more

difficult due to increased friction with the substrate; agarose beads encapsulating yeast

that do not experience the same friction because they float; and agarose blocks

encapsulating yeast, used to demonstrate microgel assembly. While validation of the

ability to culture cells was not central to this work, preliminary results in parallel

actuation and cooperative microassembly demonstrate the scalability of the microrobotic

system, and thus the potential to reach levels of processing required for tissue culture.

Overall, both the cell trapping and culture system and the bubble microrobotics

system are relatively easy to implement, do not have restrictive hardware requirements,

and are controlled by simple, self-written programs. Likewise, the microfluidic devices

used as the operation environments are also uncomplicated and require minimal

fabrication. Like similar tools used in biomedicine, extensive tests verifying cell viability

are necessary to validate the use of the techniques presented in this paper. At the same

time, the systems can also benefit by addressing the limitations encountered in practice

and conducting further characterization.

49

APPENDIX A1 Cell Handling Protocol for Trapping/Culture Device When using cells, as much of the procedure as reasonably possible should be performed in a biohood, with observance of normal sterile cell handling practices.

1. Assemble device and position on ITO heater (off) on microscope. Connect outlet

end tygon tubing to a syringe and load into syringe pump. 2. Trypsinize and retrieve cells from culture flask. 3. Transfer to test tube and centrifuge cells. 4. Remove supernatant, then wash and re-suspend pellet in PBS with gentle

pipetting. 5. Remove a small sample to check that initial cell viability is high (> 90%). 6. Remove a small sample to count cells with hemocytometer. 7. Based on the cell count determined in previous step, prepare cell suspension.

Centrifuge cells and re-suspend pellet in sufficient volume of 10% PNIPAAm solution to reach desired density. To minimize the amount of PNIPAAm solution used, determine final volume of cell-PNIPAAm solution desired, calculate how many cells are necessary to reach the desired cell concentration, and finally determine the appropriate volume of initial sample to centrifuge.

8. Fill small test tube with cell-PNIPAAm suspension. Prepare a separate test tube with each of the following: 10% PNIPAAm solution, PBS, viability assay solution. Tygon tubing extending from the device inlet will be inserted into each of the test tubes, as the solutions are needed.

9. Introduce PNIPAAm solution to prime the device using syringe pump set to negative pressure. Use negative pressure for all flushing steps to prevent the reversibly bonded device from rupturing.

10. Introduce cell-PNIPAAm solution. 11. Trap cells of interest by exposing with light. ITO heater can be turned on to

provide low heat (< 32 ˚C) to facilitate hydrogel formation. 12. Flush with PBS to remove untethered cells. 13. Switch control of heat from light pattern to ITO heater.

Note: PNIPAAm solution should be prepared at least a day in advance due to the length of time it takes for the solute to fully dissolve. The higher the concentration of PNIPAAm, the longer it will take to dissolve. Also, physical stirring is not advised, as the partially dissolved PNIPAAm will strongly adhere to the stirring instrument.

50

A2 Viability Stain Protocol for Trapping/Culture Device Protocol uses Invitrogen LIVE/DEAD® Viability/Cytotoxicity Kit for mammalian cells.

1. Prepare fresh dilution of the two dye components (Calcein AM for live cells and Ethidium homodimer-1 (EthD-1) for dead cells) in buffer solution. Ratio used (PBS:Calcein AM:EthD-1) was 100:0.5:1.

2. After flushing untethered cells from the device and transferring control to the ITO heater, introduce viability assay solution.

3. Stop flow and incubate for 20-30 minutes to allow assay to permeate the hydrogel(s). (LIVE/DEAD assay instructions note that incubation time is dependent on temperature, therefore incubation time is variable.)

4. Release the cells from the hydrogel (PNIPAAm is too opaque to image well) by lowering the temperature below LCST.

5. Use fluorescence microscopy to observe live (green) and dead (red) cells.

Note1: Calcein AM (stain for live cells) hydrolyzes and should be used within one day of preparation. Note2: Appropriate dye concentrations can vary among different cells. Optimization should be performed for each new cell type used.

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