APPARATUS FOR TISSUE ENGINEERING · circumstances, tissue engineering has emerged as a very...
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DESIGN AND DEVELOPMENT OF A BIOSTRETCH APPARATUS FOR TISSUE ENGINEERING By QIMING PANG A thesis submitted in conformity with the requirements for the degree of Doctor of Philosophy Graduate Department of Mechanical and Industrial Engineering University of Toronto ©Copyright by Qiming Pang 2009
Transcript of APPARATUS FOR TISSUE ENGINEERING · circumstances, tissue engineering has emerged as a very...
DESIGN AND DEVELOPMENT OF A BIOSTRETCH
APPARATUS FOR TISSUE ENGINEERING
By
QIMING PANG
A thesis submitted in conformity with the requirements
for the degree of Doctor of Philosophy
Graduate Department of Mechanical and Industrial Engineering
University of Toronto
©Copyright by Qiming Pang 2009
DESIGN AND DEVELOPMENT OF A BIOSTRETCH
APPARATUS FOR TISSUE ENGINEERING
DOCTOR OF PHILOSOPHY 2009
QIMING PANG
DEPARTMENT OF MECHANCIDAL AND INDUSTRIAL ENGINEERING
UNIVERSITY OF TORONTO
ABSTRACT
Tissue engineering has emerged as a promising approach to repair, replace or regenerate
damaged tissues using tissue constructs created in vitro. The standard procedure of the strategy
to create a functional tissue is to seed cells on a 3-D biodegradable and biocompatible scaffold,
to grow them under precisely controlled culture conditions provided by a bioreactor system, and
to deliver the matured construct into the patient’s body to induce and direct the growth of the
new and healthy tissue.
In this thesis, a novel bioreactor system is designed and developed, which can provide
uniaxial cyclic stretch to the tissue patch during culture process. The biostretch apparatus
employs non-contact electromagnetic force to cyclically stretch a cell-seeded three-dimensional
scaffold. The non-contact driving force and the specially designed mount allow researchers to
use standard Petri dishes and commercially available CO2 incubators to culture an engineered
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tissue patch with precisely controlled strain. The device greatly simplifies the procedure to
deliver mechanical stimulation during engineering a tissue patch.
Since the applied mechanical stimulus is generated by a magnetic force, the engineered
tissue construct is not only affected by a mechanical force, but also exposed to a magnetic field.
Thus, the effects of the time-varying magnetic field during the culture process are investigated.
The flux density of the field is modeled by COMSOL, and verified by a Gaussmeter. In
addition, one side effect of using electromagnets, that of a temperature increase, is also
investigated. The biomedical experiment results show that neither a weak low frequency
magnetic field (0.1T, 1Hz) nor an increase of 1℃ in temperature has a significant effect on the
cell culture.
The performance of the designed apparatus is verified by the biomedical experiments
from the aspects of cell proliferation and reorganization. Moreover, the mechanical parameters
(strain distribution, strain rate, and stretch force) provided by the apparatus have also been
quantitatively investigated.
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Acknowledgments
I would like to thank my supervisors, Professor Jean W. Zu and Professor Ren-ke Li, for
their guidance, support and encouragement. Their knowledge and enthusiasm gave me great
strength throughout my thesis research and have made a profound impact on my future career
endeavors.
I would like to extend my appreciation to my Ph.D. committee, Professor C.A.
Simmons, Professor W.L. Cleghorn, Professor K. Howard, Professor J. Yeow (University of
Waterloo), for their insight, suggestions and time in evaluating my research.
I would also like to thank my colleagues for their scientific discussions and friendship. I
would like to especially thank Hansong Xiao, Peyman Honarmandi, Adebukola Olatunde, and
Ming Jia. They helped to create a comfortable work environment with great enthusiasm and
creativity. Special thanks to Anson Wong for his assistance for testing, to Shuhong Li and Linda
Li for their friendly support and assistance during the years we spent together in the laboratory.
Finally, I dedicate this work to the love of my husband, Hai, for his love, patience, and
unconditional support during these years. To my son Jack who is an unending source of strength
in my life, and to my parents whose love and guidance made this work possible. I am indebted
to them. Thank you so much. I love you all!
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Table of Contents
ABSTRACT ................................................................................................................................. ii
Acknowledgments ...................................................................................................................... iv
Table of Contents ........................................................................................................................ v
List of Tables .............................................................................................................................. ix
List of Figures .............................................................................................................................. x
Nomenclature............................................................................................................................ xiii
Chapter 1...................................................................................................................................... 1
Introduction ................................................................................................................................. 1
1.1 Background ........................................................................................................................... 1
1.2 Literature Review.................................................................................................................. 4
1.2.1 Mechanical Stimulation ............................................................................................... 4
1.2.2 Electrical Stimulation................................................................................................... 5
1.2.3 Magnetic Stimulation ................................................................................................... 6
1.2.4 Bioreactors ................................................................................................................... 7
1.3 Objectives of the Thesis ........................................................................................................ 9
1.4 Thesis Overview.................................................................................................................. 10
Chapter 2.................................................................................................................................... 12
Design and Development of a Uniaxial Biostretch Apparatus.............................................. 12
2.1 Introduction ......................................................................................................................... 13
2.1.1 Types of Stretch Systems ........................................................................................... 13
2.1.2 Devices for Uniaxial Stretch ...................................................................................... 15
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2.2 Design Principle .................................................................................................................. 20
2.3 Design of Mechanical System............................................................................................. 22
2.3.1 Culture Board ............................................................................................................. 22
2.3.2 Clamps........................................................................................................................ 23
2.3.3 Mounting Tray ........................................................................................................... 24
2.4 Design of Control System ................................................................................................... 26
2.4.1 DAQ PCI-6601........................................................................................................... 26
2.4.2 Definition of the Stretch Plan..................................................................................... 28
2.5 Design of Application Software.......................................................................................... 29
2.5.1 Main Form.................................................................................................................. 30
2.5.2 Setup Form................................................................................................................. 31
2.5.3 Control Flow Chart .................................................................................................... 32
2.6 Conclusions ......................................................................................................................... 33
Chapter 3.................................................................................................................................... 35
Effect of the Time-varying Electromagnetic Field on Cell Biology...................................... 35
3.1 Introduction ......................................................................................................................... 36
3.2 Modeling the Time-varying Electromagnetic Field ............................................................ 38
3.2.1 Electromagnetic Theory ............................................................................................. 38
3.2.2 Electromagnetic Model Analysis ............................................................................... 41
3.3 Experimental Verification of the COMSOL Model............................................................ 47
3.3.1 Flux Density vs. Distance .......................................................................................... 47
3.3.2 Flux Density vs. Exciting Frequency ......................................................................... 50
3.3.3 Flux Density vs. Exciting Current.............................................................................. 51
3.3.4 Comparisons of Simulation and Experiment Results................................................. 52
3.4 Effect on Cell Biology......................................................................................................... 55
3.4.1 Cell Proliferation........................................................................................................ 55
3.4.2 Cell Morphology ........................................................................................................ 58
3.5 Heating Effect of the Electromagnet................................................................................... 60
3.5.1 Experiment Results .................................................................................................... 60
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3.5.2 Bio-experiment Verification ...................................................................................... 63
3.5.3 Effect Factors ............................................................................................................. 65
3.6 Conclusions ......................................................................................................................... 66
Chapter 4.................................................................................................................................... 68
Effect of the Uniaxial Cyclic Stretch on Cell Biology ............................................................ 68
4.1 Introduction ......................................................................................................................... 69
4.2 Validation of the Uniaxial Cyclic Stretch ........................................................................... 71
4.2.1 Cell Proliferation........................................................................................................ 71
4.2.2 Cell Orientation.......................................................................................................... 73
4.2.3 Discussion .................................................................................................................. 76
4.3 Study of Strain Distribution ................................................................................................ 77
4.3.1 Measurement Methodology ....................................................................................... 78
4.3.2 Results ........................................................................................................................ 80
4.4 Study of Strain Rate ............................................................................................................ 85
4.4.1 Measurement Methodology ....................................................................................... 85
4.4.2 Results ........................................................................................................................ 86
4.4.3 Discussion .................................................................................................................. 91
4.5 Study of Stretch Force......................................................................................................... 91
4.5.1 Electromagnetic Force Calculation ............................................................................ 92
4.5.2 Electromagnetic Model Analysis ............................................................................... 93
4.5.3 Results ........................................................................................................................ 95
4.6 Conclusions ....................................................................................................................... 100
Chapter 5.................................................................................................................................. 102
Summary and Future Research............................................................................................. 102
5.1 Summary ........................................................................................................................... 102
5.2 Future Research................................................................................................................. 103
5.2.1 Optimizing the Electromagnet Model...................................................................... 104
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5.2.2 Implementing Bio-experiments................................................................................ 104
5.2.3 Calibrating the stretch force ..................................................................................... 105
5.2.4 Integrating the mechanical and electrical stimuli together ...................................... 105
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List of Tables
Table 2.1 Uniaxial Stretch Devices for Tissue engineering…………………….… 18
Table 3.1 Nomenclature …..….……..….……..….……..….……..….…………… 39
Table 3.2 Geometrical Parameters …..….……..….……..….……..….……………42
Table 3.3 Electrical Parameters..…...……………………….…………….…….… 45
Table 3.4 Magnetic Flux Density vs. Exciting Frequency .….…………….…….…50
Table 3.5 Cell Numbers of Magnetized and Control Samples…………….…….…57
Table 3.6 Cell Numbers at 38℃ and at 37℃ ……………….…………….…….…64
Table 4.1 Cell Numbers of Stretched and Control Samples….…………….…….…72
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List of Figures
Figure 2.1 Motor-driven Uniaxial Stretch Apparatus ………………………….…… 17
Figure 2.2 Electromagnetic-driven Uniaxial Stretch Apparatus ………………….… 20
Figure 2.3 Schematic Diagram of the Biostretch Apparatus …………………….… 21
Figure 2.4 Culture Board ………………….…………………….………………… 23
Figure 2.5 Clamps ………………………….……………………….……………… 24
Figure 2.6 Mounting Tray and the Assembled Culture Dish …………………….… 25
Figure 2.7 Port Allocation of DAQ ……………………………………………….… 27
Figure 2.8 Definition of a Stretch Plan …….…………………….………………… 28
Figure 2.9 Main Form ……………………….…………………….………………… 30
Figure 2.10 Setup Form……………………….……………………….……………… 31
Figure 2.11 Control Flow Chart of the Software……………………….………………32
Figure 2.12 Assembled Bioreator System ………………………………………….… 34
Figure 2.13 Assembled Culture Dish with Electromagnets………….…………………34
Figure 3.1 Meshed Electromagnet Model ……………………………………….……43
Figure 3.2 Distribution of Magnetic Flux Density…………………….………………44
Figure 3.3 Magnetic Flux Density Bx along y Direction ………………………….… 46
Figure 3.4 Simulation Results vs. Testing Results of Magnetic Flux Density ……… 49
Figure 3.5 Exciting Current vs. Flux Density .……………………….……………… 51
Figure 3.6 Magnetic Flux Density Distribution of the Stretch Device…………….… 54
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Figure 3.7 Cell Morphology of SMC …………………………………………….……59
Figure 3.8 Cell Morphology of BMSC ……………………………….……………… 60
Figure 3.9 Temperature vs. Exciting Time under Room Temperature …………….… 61
Figure 3.10 Temperature vs. Exciting Time in incubator……………………….………62
Figure 4.1 Gross Structure of the Engineered Tissue ………………….………………74
Figure 4.2 Comparison of Cell Numbers…………………………………………….…75
Figure 4.3 Edge Structure of the Engineered Tissue ………………….……………… 76
Figure 4.4 Images of the Gelfoam and GE RTV 6166 …………………………….… 80
Figure 4.5 Strain Distribution of a Typical Sample of Gelfoam……….………………81
Figure 4.6 Distortion of a Marked Dot …………………………………………….…82
Figure 4.7 Strain Distribution for the Middle Rows of the Gelfoam and RTV Silicone
Samples ………………….………………………………………………… 83
Figure 4.8 Stretch Strain Rate at Different Stretch Strength ……………………….… 87
Figure 4.9 Retreat Strain Rate at Different Stretch Strength ……………………….…89
Figure 4.10 Stretch Frequency ………………………………………….……………… 90
Figure 4.11 Meshed Model of an Electromagnet and a Steel Bar…………………….… 94
Figure 4.12 Magnetic Flux Density, x component……………………….………………94
Figure 4.13 Magnetic Flux Density Bn along x Direction…………………………….…96
Figure 4.14 Magnetic Flux Density Bn along y Direction…………………………….…97
Figure 4.15 Comparison of Magnetic Flux Density Bx ……………………………….…97
Figure 4.16 Exciting Current vs. Magnetic Force……………………….……………… 99
Figure 4.17 Distance vs. Magnetic Force……………………………………………….100
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Figure 5.1 Schematic Diagram of the Upgraded Device ……………….……………106
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Nomenclature
English Characters
A
Magnetic vector potential
zA
Magnetic vector potential along z direction
B
Magnetic flux density
Bx Magnetic flux density, x component
By Magnetic flux density, y component
Bn Magnetic flux density, normal component
D
Electric flux density
d Thickness in z direction
E
Electric field intensity
F
Magnetic force
fext External volume force
H
Magnetic field intensity
J
Volume current density
eJ
External volume current density
zeJ
External volume current density along z direction
M
Magnetization vector
n1 Outward normal direction of material 1
n2 Outward normal direction of material 2
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P
Electric polarization vector
p Air pressure
Q Heat energy
q Free charge
Rod_force_x Stretching force along x direction
T Stress tensor
T1 Stress tensors of material 1
T2 Stress tensors of material 2
t Temperature
V Electric scalar potential
x Stretching direction
y perpendicular direction of the stretch direction
z perpendicular direction of the stretch plane
Greek Characters
Permittivity
0 Vacuum permittivity
Density
0 Vacuum permeability
r Relative permeability
Conductivity
Angular frequency
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Chapter 1
Introduction
1.1 Background
Organ failure and tissue loss are some of the most devastating and costly problems
affecting current medicine. Organ or tissue transplantation is a well known method to counter
these problems, but it remains an imperfect solution due to donor shortages and potential side
effects of immunosuppression that the recipients may have to endure lifelong [1]. Given these
circumstances, tissue engineering has emerged as a very promising approach to repair, replace
or regenerate damaged tissues using tissue constructs created in vitro.
Tissue engineering, which was formalized in the late 1980s [2], is a multidisciplinary field
combining diverse aspects of engineering, the life sciences, and clinical medicine. It contains
three general strategies to create functional tissue: isolated cells or cell substitutes, tissue-
inducing substances, and cells placed on or within matrices [1, 3]. Among these strategies, the
third strategy is in mainstream use in current tissue engineering. The standard procedure of the
strategy is to seed isolated specific cells on a 3-D biodegradable and biocompatible scaffold, to
grow them under precisely controlled culture conditions provided by a laboratory apparatus, and
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to deliver the matured construct into the patient’s body to induce and direct the growth of the
new and healthy tissue [4]. The concept of using the 3-D scaffolds has three major advantages:
(1) The 3-D scaffolds may replace the missing or damaged tissues and provide functional cell-
based substitutes of the native tissues; (2) The size, shape, and mechanical properties of the
engineered tissue patch can be controlled in vitro to treat large-scale defects or damages; (3)
Engineered tissues can provide high-fidelity in vitro models for basic studies of cell functions,
drug screening and responses to physical stimuli [5].
Although the preliminary results of tissue engineering research have helped to develop new
concepts and theories of medical technology and tissue repair, this area is still in its infancy, and
facing significant difficulties and challenges. Firstly, the artificial scaffolds have limited
diffusive ability, preventing nutrients and oxygen from being transported to cells for metabolic
activities. Therefore, acellular tissue-like constructs only have a maximum thickness of 100 µm,
or less than 10 cell layers, and mainly exist at the outer layer of the artificial scaffolds [4].
Additionally, proper mechanical properties of the tissues and organs are critical for them to
achieve their in vivo biomechanical functions, but the cell density and mechanical properties of
engineered tissue patches remain markedly poorer than those of the native tissues [5]. For
example, a lung fuctions under the respiratory pressure, and the vascular system pulses by
hemodynamic shear stress, so the magnitude and frequency of the mechanical loading that
tissues and organs are subject to in vivo can be very large. An engineered tissue patch cultured
under static environment may be unable to withstand the loading at the time of graft, which is a
known potential cause of graft failure in experimental animal studies and preclinical trials [6].
In fact, integrating cells into tissues is a very intricate process. In order to mimic the
morphological, physiological and functional properties of the native tissue, we need to advance
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our understanding of the basic principles governing tissue formation, function and failure [7].
Although D’Arcy Thompson proposed that physical forces act as causative agents during tissue
morphogenesis even in the early twentieth century [8], we have still not determined how the
physical forces affect the synthesis and functions of tissues and organs. We are equally
uncertain as to how the living systems (cells, tissues or organs) transfer the dynamical
information and regulate their networks and behaviours. To date, there are still many unknown
puzzles about intercellular interactions, such as the formation of extracellular matrix, as well as
the mechanotransduction in cells.
The above challenges cause significant research in the following areas to develop a
functional engineered tissue: the selection of cell sources, the optimization of scaffolds, and the
creation of bioreactors [1], [7]. The cell sources for implantation include autologous cells (from
a patient), allogeneic cells (from other human donors), and xenogenetic cells (from different
species). In choosing the cell sources, the issues such as ethics, safety and efficacy need to be
considered. Scaffolds may be made of natural or synthetic materials, and must be well matched
with the organ type. For the optimization of the scaffolds, there are many issues that need to be
taken into account, such as the ability for cell adhesion and the degradation time of the scaffold.
The creation of bioreactors means providing the physicochemical environment for the cells to
develop a functioning tissue. Instead of simply pumping the culture medium during culture
process, the next generation of the bioreactors combines more external stimuli during tissue
culture. However, the issues about cell sources and scaffold optimization will not be addressed
further in this thesis. The major objective of the thesis is to design and develop a novel
bioreactor system, which can provide mechanical stimulation (uniaxial stretch) to the tissue
patch during culture process.
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1.2 Literature Review
Culturing a tissue patch in a dynamic in vitro microenvironment is important in guiding the
formation of tissue with certain structural and functional characteristics [1], [7]. Scientists
believe that external guides and signals, such as mechanical stress and strain, magnetic, and
electrical stimuli, are important to help cells to grow into functional 3-D implantable tissues [7,
9-11]. In 1988, Vandenburgh et al. [12] and Terracio et al. [13] imposed cyclic stretch on
skeletal and cardiac myocytes, and observed that mechanical stimulation is crucial for the
differentiation and orientation of tissue development. Meanwhile, Radisic et al. [14] found that
electrical field stimulation induced cell alignment and coupling. In this section, a brief literature
review is carried out on the aspects of applying external physical stimuli for tissue engineering.
1.2.1 Mechanical Stimulation
Providing appropriate physiological stresses and strains may help to form correct molecular
and macroscopic architecture during engineering the tissue in vitro, which is crucial for
obtaining proper tissue functions [7, 11]. Therefore, as long as formulating the concept of tissue
engineering, mechanical stimulation has been considered. In 2000, Brown classified the
techniques for mechanical stimulation of cells according to the primary loading modality:
compression loading, longitudinal stretch, bending, axisymmetric substrate bulge, in-plane
substrate distention, and fluid shear stress, etc. [15]. Among these techniques, three loading
modes, compression, stretching, and fluid flow, are the most commonly used methods.
Mechanical stimulation has been shown to increase cellular alignment, proliferation, gene
expression and construct morphology in many different cell types [11, 12, 16]. However,
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previous studies have demonstrated that cells throughout the human body are exposed to
various forms of mechanical stimuli, which means that different tissues are sensitive to different
mechanical stimuli [17].
Chondrocytes, cells found in cartilage, are sensitive to compressive loading since cartilage
tissues are primarily subjected to this kind of loading in vivo. Compression has also been
utilized for other cell/tissue types such as bone [18-20]. Other cell types, such as endothelial
cells in blood vessel and osteocytes in bone tissue, are subject to shear stress in vivo. Therefore,
fluid flow is often successfully used as the mechanical stimulation in this kind of cell culture
[21-23]. However, ligaments, muscles such as skeletal muscle, cardiac muscle and smooth
muscle are usually under cyclic stretch in vivo. For these types of cells, cyclic stretch is served
as an important regulator [6, 12, 16, 24]. The types of stretching devices and the effects of cell
biology will be reviewed in Chapter 2 introduction section.
Although literature shows that mechanical stimulation is an important regulator for the
development of engineered tissues with similar construction and mechanical properties as native
tissues, significant work is needed to identify the types of mechanical stimulation required to
optimize the formation of a functional engineered tissue. To date, it is still unclear how the
mechanical force is transferred to the individual cells, and what kinds of responses are yielded
at the molecular and cellular level. Therefore, quantitative study of the influence of mechanical
force is necessary for functional tissue engineering.
1.2.2 Electrical Stimulation
Many tissues and organs in vivo are not only subjected to mechanical loading in order to
achieve their functions, but also exposed to some electrical signals. For instance, the nervous
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system utilizes neurons and nerve cells to build an electrochemical wiring network, and to
generate and transform the electrochemical impulses. In addition, the contraction of
myocardium (cardiac muscle) is controlled by an electrical signal generated by the sinoatrial
node. These facts suggest that electrical signals can also be used as an external factor to
stimulate the creation of an engineered tissue patch [25].
Previous studies show that when a weak DC electric field was applied to embryonic
fibroblasts cultivated on a 2-D substitute, the fibroblasts became oriented perpendicular to the
electric field lines and migrated towards the cathodal end of the field [26]. The group of Cho et
al. [27] explored the cellular behaviours on the 3-D matrix with the stimulation of a DC electric
field, and demonstrated that fibroblasts were more readily reoriented than rat mesenchymal
stem cells. Radisic et al. [14] reported that a pulsed electrical field stimulus induced cell
alignment and coupling, improved cell contractile function, and resulted in a remarkable level
of ultra-structural organization. They also found that electrical field stimulation can enhance
cell elongation and orientation, which is directly related to the functional properties of the
engineered contractile tissues [28].
1.2.3 Magnetic Stimulation
To facilitate the development of engineered tissue, besides mechanical and electrical
stimulations, researchers are also seeking for other stimulations. Magnetic field (MF) has been
used as a simple and effective means to orient collagen gels, since collagen gels provide the
matrix substrate circumstance which directly influences the cell growth. In 1993, Tranquillo et
al. [29, 30] presented that a strong magnetic field (Tesla order) can be used to orient collagen
fibrils and subsequently orient the entrapped smooth muscle cells. Hashimoto et al. [31]
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demonstrated that moderate magnetic field (millitesla order) can enhance cell migration,
depending on the types of cells. Even a weak electromagnetic field (Gauss order) with low
frequency (<300Hz) has been reported to have an effect on enzyme reaction and transcript
levels for specific genes [32]. Thus, magnetic stimulation can positively affect cell growth in a
wide range of the MF magnitude.
Although studies show that both magnetic and electrical fields can influence the tissue
culture process to some degree compared with purely mechanical stimulation, reports of in vitro
culturing under electric or magnetic stimulus are still rare. Thus, intensive investigation is
needed to carry out to study the influence of physical stimuli, such as mechanical, electrical and
magnetic, on cell morphology and behaviours that are important for controlling engineered
tissue constructs.
1.2.4 Bioreactors
During tissue culture process, the use of scaffolds and laboratory apparatus has been
regarded as a crucial technique. Although there is no universally recognized definition of these
laboratory apparatus, they can be called bioreactors. Bioreactors, with respect to Freed et al.
[33], can be defined as “laboratory tissue-culture devices, which provide a controllable,
mechanically active environment that can be used to study and potentially improve engineered
tissue structure, properties, and integration.” Even though bioreactors should be designed and
fabricated following specifications that differ from tissue to tissue, cell-culture parameters such
as temperature, pH, biochemical gradients, and mechanical stimulation must be continuously
controlled during the maturation period because bioreactors should provide an in vitro
environment mimicking the in vivo conditions [34, 35]. Thereby, the design of an appropriate
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bioreactor for a specific tissue is important, but also difficult in that bioreactors provide all the
environment control and regulatory factors necessary for cell culture. Compared with
conventional static culture systems, modern bioreactors normally integrate perfusion system
and/or external physical stimuli in order to regenerate a functional engineered tissue.
To improve nutrient and oxygen transportation, researchers have designed several
bioreactors with different perfusion systems [36], [37] [38]. A basic fluid-dynamic cultivation
vessel is the spinner flask that provides a well-mixed environment around cells. The engineered
tissues growing in a spinner flask have better metabolic activity and morphological appearance
than under static conditions [36]. However, the spinner flask may not be the optimal cultivation
vessel for cells. It causes the turbulent fluid flow at the surface of the constructs. The turbulent
fluid is usually characterized by eddies which can destroy the seeded cells [4, 39]. Fluid-
dynamic cultivation environment can only partially compensate the absence of capillaries by
providing nutrients and gas through the entire thickness [36, 40], meaning that merely adding
perfusion system in the bioreactor is not a powerful way to increase mass diffusion.
Some research groups consider integrating multi stimuli in the culture process. The most
common method is to combining the perfusion system and mechanical stimuli together in one
bioreactor. For example, a continuous pulsatile perfusion system is integrated with mechanical
stimuli for vascular tissue [41, 42]; the perfusion system and static/dynamic compression
loading are applied to cartilage tissue [43], and the perfusion flow and fluid flow-induced shear
stress are combined in bone tissue engineering [44]. Moreover, some groups considered
integrating more than one external stimulus into the system. For instance, Feng et al. [45]
developed a device to provide both electrical stimulus and dynamic tensile force during culture
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process. This device tries to solve the coordination problem between the dynamic mechanical
loading and the spontaneous beating.
However, the above techniques have a number of shortcomings that limit their applicability
for in vitro applications. First of all, the relationships between external stimuli and cell reactions
are still unclear. Secondly, using most of the above bioreactors to make tissue constructs
requires a complex casting process beyond commercially available standard cell culture dishes
and CO2 incubators. These specifically designed apparatus make the size and quality of tissue
patch variable and the culture process difficult to operate. In addition, applying external
stimulation during the culture process increases the risk of contamination, especially for long-
term in vitro culture, due to the physical connection and the complicated assembly procedure.
Therefore, design a simple bioreactor system is essential for researchers to further their study.
1.3 Objectives of the Thesis
The main objective of the thesis is to design a bioreactor system that can provide
mechanical stimulation under standard culture frames. It includes a biostretch apparatus that can
provide uniaxial cyclic stretch to a three-dimensional cell-seeded scaffold. The apparatus
employs non-contact electromagnetic force as driving force. The non-contact driving force,
together with the specially designed mount, allows researchers to use standard Petri dishes and
commercially available CO2 incubators to culture an engineered tissue patch under well defined
mechanical conditions. Based on this bioreactor system, the effects of the magnetic field and
uniaxial stretch on cell biology have been studied.
The prototype of the apparatus has been utilized in a cardiac tissue research lab at the
University Health Network in Toronto. Although the apparatus is initially designed for cardiac
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tissue engineering research, the apparatus has the potential to be used in other tissue engineering
areas such as bone and musculoskeletal tissue. The objectives are pursued through the following
sub-objectives:
Developing a prototype of the apparatus, including the mechanical system, computer-
based control system and application software
Modeling the electromagnetic field, and investigating the magnetic effects on cell
proliferation and morphology
Investigating the effects of uniaxial stretch from the aspects of cell proliferation and
organization, and quantitatively studying the mechanical parameters provided by the
apparatus, such as strain distribution, strain rate and stretch force
1.4 Thesis Overview
The main body of this thesis consists of three chapters, which discuss three tasks mentioned
in the previous section. They are closely concatenated and organized as a whole system.
However, they can also be treated as three different studies in the related corresponding areas.
In Chapter 2, Design and Development of a Uniaxial Biostretch Apparatus, the past uniaxial
stretch devices are first reviewed. The development procedure of the biostretch apparatus is
elaborated from the aspects of the design principles, the mechanical system, the control system,
and the application software.
In Chapter 3, Effect of the Time-varying Electromagnetic Field, A time-varying
electromagnetic field is studied, since the apparatus uses non-contact force as stretch force. The
field is modeled by commercial software COMSOL. The magnetic flux density B and field
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distribution are simulated and verified. The effect of the time-varying magnetic field with
temperature fluctuation on cell proliferation and morphology are investigated.
In Chapter 4, Effect of the Uniaxial Stretch, the performance of the apparatus has been
tested by the bio-experiments from the aspects of cell proliferation and organization. The
mechanical parameters (strain distribution, strain rate, stretch force) provided by the device
have been investigated.
In Chapter 5, Summary and future research, the results obtained in this research are
summarized and the future research directions in this area are proposed.
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Chapter 2
Design and Development of a Uniaxial
Biostretch Apparatus
The objective of this chapter is to design and develop a prototype of the uniaxial biostretch
apparatus for tissue engineering research, especially for cardiac engineered tissue. The
biostretch apparatus employs non-contact electromagnetic force to cyclically stretch a cell-
seeded three-dimensional scaffold. The non-contact driving force and the specially designed
mount allow researchers to use standard Petri dishes and commercially available CO2
incubators to culture an engineered tissue patch with precisely controlled strain. The device
greatly simplifies the procedure to deliver mechanical stimulation during engineering a tissue
patch. This chapter mainly describes the design principles including the mechanical system,
control system and application software.
In this chapter, a brief introduction of the currently used stretch apparatus is presented in
section 2.1. Section 2.2 describes the design principle and criteria; section 2.3, 2.4, and 2.5
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introduce the design of mechanical system, control system and application software,
respectively. Concluding remarks are drawn in section 2.6.
2.1 Introduction
It has long been recognized that mechanical signals can improve the organization,
composition, and function of engineered tissue, such that mechanical stimulation is the most
common physical stimulation in a bioreactor system design. These mechanical stimuli can
successfully mimic the in vivo activities of daily living. However, before utilizing these
mechanical signals to “precondition” the tissue-engineered constructs, the magnitudes of the
mechanical parameters in vivo have to be estimated. Literature review shows that the strain of
tendon tissue can reach 2.4% [46]. Meanwhile, previous research demonstrated that the
circumferential strain from a complete set of the planar tagged images of a human left ventricle
ranges from 25% shortening to 10% stretching, while the strain of a paced canine left ventricle
can range from 20% shortening to 20% stretching [47].
Although mechanical stimuli can be categorized into different groups, as we discussed in
chapter one, only three major loading modes, compression, stretching, and fluid flow are
commonly used [15], [17] . Since my initial study object, cardiac tissue, along with
musculoskeletal and vessel tissues, are all sensitive to stretching, the classification and
application of stretch systems are emphasized here.
2.1.1 Types of Stretch Systems
There are two major stretch methods for cell culture, uniaxial stretch and biaxial stretch. In
biaxial stretch devices, cells are usually cultured on a circular membrane and the force is
13
commonly generated by pneumatic pressure or pistons [48-52], such that the membrane
deforms in both the radical and circumferential directions. These systems have been used for
culturing fibroblast, tendon, and annulus cells. One of the pneumatic driven stretch devices is
commercialized, called Flexcell® culture systems. However, the strain profile of the biaxial
stretch device [48, 50] varies, with the highest strain along the culture walls and lowest strain in
the center. These variations cause difficulties in analysis the relationship between the strain and
cell effects.
Uniaxial stretch is another commonly used stimulus, because it can offer a very precise
control of end-to-end displacement that is important to evaluate the average strain of the
scaffold [53]. This type of stimulation provides a versatile platform for researchers to
extensively study cell response and its mechanical properties. Uniaxial cyclic stretch exhibits
improvements of morphology and organization of engineered tissues by the following groups.
The group of Vandeburgh [12], [54][55] indicated that mechanical stimulation affected protein
accumulation and localization, while mechanical stimulation also improved cellular
proliferation, myofiber organization. The in vitro generated artificial muscle organs contain
parallel networks of long unbranched myofibers organized into fascicle-like structures.
Eschenhagen et al. and Zimmermanne el al. [2][16][24] presented that chronic mechanical
stretch of engineered heart tissue improved organization of cardiac myocytes into parallel arrays
of rod-shaped cells, and increased cell length and width, myofilament length, and mitochondrial
density. The group of Terracio [13] also imposed mechanical stretch on particular cell types
(cardiac myoctyes, endothelial cells, or fibroblasts) and showed that mechanical stimulation had
a favorable effect on cell orientation. Li et al. [6] cultured isolated heart cells on gelatin-matrix
scaffolds, and demonstrated that cyclical mechanical stretch improved the proliferation and
14
distribution of the seeded human heart cells, and stimulated the formation and organization of
extracellular matrix, which contributed to the improvement in the mechanical strength of the
cardiac graft.
2.1.2 Devices for Uniaxial Stretch
According to the methods of driving force application, the uniaxial stretch systems can be
classified mainly into two groups, motor-driven and magnetic force driven, shown in Table 2.1.
Motor-driven Devices
The first group uses step-motors to provide tension. Vandeburgh et al. [54-56] [12] started
studying mechanical stimulation in the 1980s, and developed a mechanical cell stimulation
device (version four) shown in Fig. 2.1 (a). The cell-seeded construct is combined with two
steel pins. One pin is immobile, and the other pin attaches to a step-motor (SM), which applies
direct mechanical stretch. This device can hold six plates, and apply sensitive force transducers
(LC) to detect the loading force. With this device, users can define uniaxial the amplitude,
velocity, and pattern of the stretch. Eschenhagen et al. and Zimmermann et al. [16, 24][57],
developed a uniaxial cyclic stretch apparatus driven by a stepper motor, shown in Fig. 2.1 (b).
This device provided a circular mold, and phasic mechanical stretch. It was reported that the
circular geometry causes a homogeneous force distribution throughout the tissue. This phasic
stretch induced hypertrophic growth and marked functional improvement. Meanwhile, the
group of Terracio also designed motor-driven cell stretch devices to study the cardiac cell
behaviors [13, 58], shown in Fig. 2.1 (c). This device has a dual-stretch unit. The unit consists
of a TiN-coated stainless steel yoke (C) assembled with dual-rod sliders (E). One side of the
15
silicone stretch membrane is clamped with the slider, and the other side is combined with a poly
amide-imide base plate (F). The stepper motor (A) drove the yoke through lead screw (B). The
stretching was processed inside the culture dish (D). Another type of motor-drive device is
shown in Fig. 2.1 (d) [59, 60]. This device consisted of a computer equipped with a four-axis
motion control card, customize software, and four stepper motor controllers operating four
independent cycliuc strain units.
These motor-driven devices allow great versatility of input waveform, which provides
researchers more flexibility to define the stretch patterns. However, in order to apply cyclic
stretch, the motor and the engineered tissue constructs must be in physical contact, which
requires both the culture dishes and the bioreactor system to be customized. Therefore, the
assembly procedure of the culture systems is very complex, which may increase the risk of
contamination and the variability of the products.
16
(a) (b)
(c) (d)
Figure 2.1 Motor-driven Uniaxial Stretch Apparatus
17
Table 2.1 Uniaxial Stretch Devices for Tissue Engineering
Authors Cell Source Methods of Force Application Parameters studied
Terracio et al. [13, 58] Rat cardiac fibroblast Stepper-motor driven strain
Vandeburgh et al. [54-56] [12] Skeletal muscle myoblast Stepper-motor driven Strain, force
Elson et al.[61, 62] Cardiac fibroblasts,
chicken embryo fibroblasts
Stepper-motor driven Strain, force, stiffness,
frequency
Eschenhagen et al. and
Zimmermann et al. [16, 24][57]
Cardiac myocytes Stepper-motor driven Force, strain
Langelier et al.[63] Fibroblasts Stepper-motor driven strain
Joshi et al.[59] Human dermal fibroblasts Stepper-motor driven strain
Mol et al. [64] Human venous myofibroblasts Linear actuator driven strain
Pfister et al. [65] Rat neuroblastoma and glioma cells Voice coil actuator driven Strain, strain rate
Liu et al. [66][6] Rat aorta endothelial cells, cardiac
myocytes, smooth muscle cells
Magnetic force driven strain, deformation
Smalt et al. [67] Osteoblastic cells Magnetic force driven strain
18
Electromagnetic Force Driven Devices
The second group of uniaxial stretch systems uses electromagnetic force. In one design, a
specially designed bio-stretch apparatus with square culture dishes, shown in Fig. 2.2(a), is used
to impose the uniaxial cyclic stretch on the tissue samples. One side of the tissue patch is fixed
on the dish (lower side of the figure), and the opposite side is attached to a magnetic stainless
steel bar (upper side of the figure). The culture dish is placed in front of a programmable
electromagnet, which can provide a dynamically changing magnetic field. Stretch patterns of
variable frequency, duty cycle, and amplitude can be controlled by the magnetic field [66].
Similarly, Smalt et al. [67], designed another kind of customized culture well, which can apply
cyclic strain (500-5000µƐ, 1Hz) to a polystyrene film by using electromagnetic force, shown in
Fig. 2.2(b).
Literature review shows that uniaxial cyclic stretch is essential for the reorganization,
differentiation, and the mechanical properties of different types of tissues, and that the
magnitude and the type of these stresses/strains are specific to cultured tissues. Among these
stretch devices, I prefer the electromagnet-driven apparatus developed by Liu et al. [66]. This
apparatus used non-contact magnetic force as driving force, which allows the culture dishes to
be easily isolated from the environment and thus greatly reduces the contamination risk. With
this apparatus, Liu et al. and Yang et al. [68, 69], [70] investigated the effect of mechanical
stretch on cell proliferation and differentiation; Lewis et al. [71] studied the adaptation of
craniofacial skeletal muscle following different mechanostimulation – rapid ramp stretch or
cyclical ramp stretch with 7.5% and 15% strain; and Freed et al. [72] applied this equipment to
study the mechanical properties of hybrid cardiac constructs made from rat heart cells, fibrin,
and biodegradable elastomeric knitted fabric. Finally, the group of Li et al. [6] used this
19
apparatus to grow cardiac muscle grafts and found that cyclical mechanical stretch can improve
the mechanical strength of the engineered cardiac grafts.
(a) (b)
Figure 2.2 Eletromagnetic-driven Uniaxial Stretch Apparatus
2.2 Design Principle
The most important advantage of the electromagnet-driven apparatus is the use of non-
contact driving force, which allows researchers to culture the engineered tissue patch in an
isolated environment. However, some limitations of this apparatus [66, 67] hamper its further
utilization. First, the apparatus has a relatively complicated control system that includes a
biostretch manager, a biostretch controller, and a set of magnet boards with customized Petri
dishes. As such, the complicated assembly procedure and customized culture dishes increases
the risk of contamination. Moreover, the deformation of the tissue patch is only controlled by
magnitude of the stretching force, and a precise and controllable strain magnitude is difficult to
obtain by controlling only the electromagnetic force. One of the goals of my thesis study is to
20
design a uniaxial stretch apparatus, which can apply precisely controlled strain to the tissue
constructs with the minimum risk of contamination during a long-term tissue culture process.
The new apparatus is initially designed to culture cardiac tissue patches in a standard CO2
incubator and impose controllable uniaxial strain on the cell-seeded scaffold during the culture
procedure. However, if the apparatus can provide a wide range of stretch patterns with different
strain magnitude and frequency, it can become a yielding and versatile tool for researchers to
study different tissue cultures. The desired design criteria of the new biostretch apparatus was
summarized as follows:
It can use standard Petri dishes.
It can provide a wide range of stretch patterns with precisely controlled strain
magnitude and frequency applied on the scaffold.
It can reduce the potential for bacterial or fungal contamination during the culture
procedure.
It can be easily installed in and removed from a CO2 incubator.
Figure 2.3 Schematic Diagram of the Biostretch Apparatus
(1 - Culture boards, 2 - Incubator)
21
The apparatus consists of two parts, as shown in Fig. 2.3. The first part is a computerized
control system used outside of the incubator, and the other part is composed of up to three
culture boards used inside the incubator. Each culture board can hold three standard Petri
dishes. A cell-seeded scaffold with two stainless steel clamps at each end is mounted in the
center of the culture dish. On the two opposite sides of each culture dish are two identical
electromagnets providing forces equal in magnitude and opposite in direction to two ends of
each culture patch. Therefore, when a current is passed through both magnets, force is applied
symmetrically to the tissue patch.
2.3 Design of Mechanical System
2.3.1 Culture Board
The culture board is designed to hold culture dishes and electromagnets, and can be easily
installed in or removed from the incubator, shown in Fig. 2.4. In order to stretch a useful
number of samples in the commercial CO2 incubator, while minimizing interactions between
electromagnets, each culture board is designed to hold three dishes. One culture board consists
of two plates, the base plate (1) and the hold plate (4). The base plate is used to mount the
electromagnets and the hold plate. The hold plate is mounted in the middle of base plate, and is
used to hold the Petri dishes. Each culture board has two different sizes of hold plates that are
designed for the 35-mm and 60-mm Petri dishes, respectively. There are three pairs of V-shape
blocks (3) on the hold plate, which are used to mount the dish in the middle of two opposite
electromagnets. Combined with hold plate, the slots (2) on the base plate can be used to adjust
the distances between the culture dishes and electromagnets.
22
Figure 2.4 Culture Board
(1-base plate, 2-slot, 3-V-shape block, 4-hold plate)
2.3.2 Clamps
Since the apparatus uses non-contact magnetic force as stretch force, two stainless steel
magnetic clamps need to attach to the two ends of the scaffold, one at each end. When a
magnetic field is applied to the clamps, the clamps move under the magnetic force, causing the
scaffold to stretch. Before designing a proper clamp, two elements have to be taken into
account: Firstly, the clamps must be very simple to assemble because the combination of the
clamps and the cell-seeded scaffold has to be finished in the hood. Secondly, the clamps should
have a symmetric structure such that the tissue patch can undergo a symmetric force.
Fig. 2.5 shows the design of the clamps. The clamp (2) is made of polycarbonate, which
can withstand an alcohol disinfectant without cracking, and has some elastic property for the
stainless magnetic bar (1) to be inserted. The pressure enforces the scaffold (3), steel bar (1),
and the clamp (2) to combine together tightly. The material of the bar is Ph-174, a type of steel
that satisfies two characters – magnetic and stainless, when put into medium. The shape of the
steel bar is rectangular, so it is simple to manufacture and assemble. For a 35-mm Petri dish, the
dimensions of the steel bar are 16mm x 3.5mm x 1mm. In the hood, the operator can simply use
23
tweezers to insert the rectangular-shaped steel bar into the slot of the clamp and tighten the
scaffold.
Figure 2.5 Clamps
(1-steel bar, 2-clamp, 3-scaffold)
2.3.3 Mounting Tray
To insert the assembled tissue construct (the scaffold with clamps) into the culture dish and
to prevent motion other than uniaxial stretch, a mounting tray is used, shown in Fig. 2.6 on the
bottom left. The center groove (2) confines the scaffold to the centre of the dish, while eight
small slots (3) increase the bottom area of the scaffold that is exposed to the culture medium. To
control the deformation of the scaffold, stopper pins (1) of various sizes are selectively inserted
into the holes on each side of the mounting tray (4). The inner stopper pins (10) ensure that the
scaffold returns to its original position after stretch; the outer stopper pins (11) limit the amount
of stretch. The position and distance between the pins determine the deformation of the scaffold
24
and the stretch methods. After the tissue construct mounted on the mounting tray, the
assembled tray can be fit into the standard culture dish, shown in Fig. 2.6 right side.
Compared to the traditional stretch method (fixing one side and stretching the other side),
this device can provide two kinds of stretch methods, symmetric stretch, and asymmetric
stretch, based on the assembly of the tray. When the two clamps have same distance L
between the inner and outer stopper pins, the scaffold exhibits symmetric stretch; when the
inner and outer stopper pins tightly secure one of the clamps, and only one clamp is movable,
the scaffold performs asymmetric stretch (traditional method).
Figure 2.6 Mounting Tray and the Assembled Culture Dish
(1-stopper pins; 2-center groove; 3-slot; 4-hole; 5-clamp; 6-magnet; 7-culture dish;
8-scaffold; 9-mounting tray; 10-inner stopper; 11-outer stopper)
25
2.4 Design of Control System
The apparatus is designed to allow two different types of stretch patterns during the tissue
culture process: continuous cyclic stretch and intermittent cyclic stretch [66]. Continuous cyclic
stretch means cyclically stretching the scaffold at a constant frequency without stopping;
intermittent cyclic stretch means alternately stretching and resting the scaffold. For example, the
scaffold first experiences a period of cyclic stretch on the order of 15 minutes, called burst time,
and then sits in a static environment for a period on the order of 45 minutes, called rest time, to
avoid cell injury. The stretch plan parameters, with specific burst and rest times, are determined
by researchers. In order to conduct a designed stretch plan, a controlled current pulse train is
needed to work as the power supply of electromagnets. This pulse train contains all the
information of the stretch plan. Thus, we need to implement pulse train generators and a user-
friendly interface to customize the stretch parameters of a stretch plan.
2.4.1 DAQ PCI-6601
To fulfill the design criteria, a National Instruments data acquisition card (DAQ) NI PCI-
6601 is employed for data acquisition and pulse train generation. This DAQ card has four
up/down 32-bit Counters/Timers with a maximum source frequency of 20MHz, and up to 32
digital I/O lines. The counters can be used as the pulse train generators, while the digital I/O
lines can be used as input/output control signal channels. To obtain three entirely independent
stretch plans for three culture boards, the four stretch parameters of each stretch plan, strength,
frequency, pattern, and duration of the stretch time, must be controlled separately. Thus, three
counters of the DAQ are employed to generate three pulse trains so that the frequency and the
pattern of each pulse train can be controlled separately. The pattern means the proportion of
26
high-voltage time (ON-time) and low-voltage time (OFF-time) of a pulse period. Both the
pattern and the frequency can be controlled by setting the configuration of the counter. The port
allocation of PCI-6601 is shown as in Fig. 2.7.
For each stretch plan, the PCI-6601 needs to allocate one output channel from a counter to
generate the pulse train with a specific frequency and pattern; and another eight output channels
to an 8-bit D/A converter to control the voltage of the pulse train, so that the stretch strength can
be controlled with the accuracy of 1/128. The pulse train generated by the counter modulates
the frequency of the D/A converter’s output. After amplification, the modulated pulse train
eventually serves as the power supply of the electromagnet. For security reasons, the system
also needs four input ports to check the status of the power supply and test the connections
between the culture boards (inside the incubator) and the control system (outside the incubator).
To build up this control system, 33 out of 68 ports of the PCI-6601 have been used.
Figure 2.7 Port Allocation of DAQ
27
2.4.2 Definition of the Stretch Plan
There are different methods to add mechanical stimulation during a tissue culture process.
Researchers can design a specific stretch plan according to the specific requirement of a tissue
culture process. For instance, in the first 24 hours, we may let the cells grow in a static
environment without any mechanical stimulation. This period of time is necessary for the cells
to recover from isolation procedure and enable attachment and spreading on the scaffold. In the
next 24 hours, we may add mechanical stimulation with low frequency and strength. After
several days of light stimulation, we may gradually increase the stretch frequency and strength.
Additionally, different parameters may be used for each of the three stretch boards, allowing
researchers to study the influence of specific parameters.
Figure 2.8 Definition of a Stretch Plan
28
Fig. 2.8 shows a detailed stretch plan. A stretch plan can have several durations of stretch
with different stretch parameters (sessions). It can satisfy different stretch patterns, such as
continuous cyclic stretch and intermittent cyclic stretch. The number of the sessions and the
proportion of the burst time and rest time in a session are determined purely by the software.
Each session has its own stretch pattern. Each pattern includes two parts, burst time (stretched
time) and rest time (static time). For a continuous cyclic stretch, the rest time is set to be zero.
In each burst time, three parameters of the pulse train need to be defined: voltage (stretch
strength), frequency (stretch frequency), and ON time (duty cycle). In a period, ON time
controls the stretch holding time of a period; OFF time determines the relaxed state between
two periods. The parameters are set in the Setup Form of the application software, and fulfilled
according to the above port allocation. A stretch pattern can be performed once or
continuously.
2.5 Design of Application Software
The function of the control software is to provide a user-friendly interface where all of the
stretch parameters can be characterized and monitored. As an improvement over Liu’s
apparatus [66], the new apparatus does not limit the number of sessions for each stretch plan. It
gives researchers more flexibility to perform long-term (>2 days) study of tissue culture, which
may need several different stretch sessions. In addition, each stretch plan can have its own
stretch strength instead of three stretch plans sharing the same stretch strength. The application
software has two major forms: Setup Form and Main Form. Users design a stretch plan and
characterize each parameter from the Setup Form, however, trigger, monitor, or stop the
specifically defined stretch plans from Main Form.
29
2.5.1 Main Form
Main form is the main user interface of the apparatus. It monitors the status of the three
stretch boards. The form is divided into three sections, one for each of the stretch boards, shown
in Fig. 2.9. The left side of each board shows the stretch information, including Sessions, Status,
Strength, ON Time, Burst Frequency, and Cycle Type. Although the information is displayed
here, it cannot be modified from this form, since the Setup Form is the only place to define the
parameters. The right side of the board shows the stretching time and graph with specific
strength and frequency. At the right side of the form, there are four buttons: Run Board1, Run
Board2, Run Board3, and Run All. These buttons are used to trigger each board’s stretch plan.
The three boards can be triggered separately or simultaneously. After starting a stretch plan, the
button’s function changes to Stop Board, so that the four buttons will change to: Stop Board1,
Stop Board2, Stop Board3 and Stop All. The operator can stop any of the boards by clicking the
relevant button.
Figure 2.9 Main Form
30
2.5.2 Setup Form
The Setup Form is the only entry point where users can modify the stretch parameters. There
are three major list groups in the form: Plan List, Plan Details, and Select Plans. In the Plan List
group, several defined plans are listed. Users can add, delete or modify the plan using the
control bar underneath. In the Plan Details list group, the details of the currently selected Plan
are shown. For example, in Fig. 2.9, the currently selected plan number is 1, and the Plan
Details group lists all the stretch sessions of that plan. Users can set the parameters such as
BurstTime, RestTime, BurstFrequency, Strength, and ONTime in the group. In the Select Plans
list group, users determine the stretch plan for each stretch board by selecting the defined
stretch plans. The Setup Form gives users the maximum flexibility to design a stretch plan with
the specific stretch time, pattern, frequency and strength.
Figure 2.10 Setup Form
31
2.5.3 Control Flow Chart
Fig. 2.11 demonstrates the control flow of the software. When the application software is
activated, the control thread starts working. First, it controls three boards to load their own
stretch plan. Then periodicly checks the statuses of three boards, Board1, Board2, and Board3.
When the thread finds any of the boards need triggering, it will control the corresponding board
to perform burst or rest action according to the defined stretch plan. When more than two
objects are triggered, the control thread will determine the next action according to its own
timeline. The major function of the control thread is coordinating the action of each stretch plan.
Figure 2.11 Control Flow Chart of the Software
32
2.6 Conclusions
This chapter describes the development of a uniaxial biostretch apparatus to culture
engineered tissue patches under well defined mechanical conditions. The apparatus provides a
versatile platform to study the influence of the mechanical stimuli on cell culture. The new
apparatus has demonstrated several technical and practical advantages to existing ones:
It allows tissue patches to be cultured in a standard culture dish and provides two
kinds of stretch methods, asymmetric stretch and symmetric stretch.
It can deliver a wide range of precision consistent strains and stretch frequencies.
The use of uniquely designed scaffold mounting trays, with non-contact force, allows
researchers to use standard Petri dishes, which significantly simplifies the assembly
process, and ultimately reduces the potential for contamination.
The device is easy to sterilize and clean, which makes delivering mechanical
stimulation easy and simple.
With this device, various sizes of culture tissues (35 and 60 dishes) are available.
A prototype of the apparatus, shown in Fig. 2.12, was constructed and tested in a research
lab at the University Health Network in Toronto. Fig. 2.13 shows the assembled culture dish
mounted between a pair of electromagnets.
Since the device utilizes non-contact magnetic force as its driving force, the effect of the
magnetic field will be studied in the following chapter. The calibration of the device and the
effects of mechanical stimulus will be discussed in Chapter 4.
33
Figure 2.12 Assembled Bioreactor System
Electromagnet Culture Dish Gelfoam Stopper Steel Bar
Figure2.13 Assembled Culture Dish with Electromagets
34
Chapter 3
Effect of the Time-varying
Electromagnetic Field on Cell Biology
The objective of this chapter is to study the effect of the time-varying electromagnetic field
generated by a horseshoe-shaped electromagnet, which is utilized in the stretch device described
in Chapter 2. The magnitude and distribution of the electromagnetic field is modeled and
simulated with COMSOL Multiphysics 3.3, which also provides an effective tool to
quantitatively study the applied magnetic force in Chapter 4. The magnetic field influence on
cell biology is investigated from the aspects of cell proliferation and morphology.
In this chapter, section 3.1 briefly introduces the current state of the understanding of
magnetic field influence on cell biology. Section 3.2 presents the modeling procedure of the
time-varying electromagnetic field. Section 3.3 verifies the model through physical
experimentation. Section 3.4 describes the study of magnetic influence on cell biology through
two bio-experiments. These experiments are performed in a research lab at the University
35
Health Network in Toronto. Section 3.5 illustrates the heating effects of the electromagnet on
cell proliferation. Conclusions are discussed in section 3.6.
3.1 Introduction
The uniaxial cyclic stretch apparatus that was described in Chapter 2 uses electromagnetic
force as the driving force. The primary advantage of magnetic force is that it can directly apply
controlled force on an engineered tissue patch without introducing a pathway of infection.
Besides acting as the driving force to stretch an engineered tissue patch, the magnetic force has
also been used in other recent studies along with magnetite nanoparticles for cell seeding [73],
gene transfection [74], and constructing multilayered cell sheets with heterotypic cocultured
cells [75] in tissue engineering. This promising technique developed a new branch of
engineered tissue called magnetic force-based tissue engineering [75], [76, 77]. However, the
biological effects of the magnetic field during tissue culture have not yet been fully
investigated.
The most enigmatic external source of stimulation in tissue engineering is the magnetic
field, since all biological systems are exposed to magnetic fields from the earth or other sources.
Institutions such as EMF-RAPID (Electric and Magnetic Fields Research and Public
Information Dissemination) and NIEHS (National Institute of Environmental Health Sciences)
started to study the influence of magnetic fields on the human body in 1992 [78]. It was
reported that although magnetic fields maybe have a negative effect on the human body, that
effect, if any, is considerably small [31]. To date, the effects of magnetic fields to human beings
have been continually studied in a wide range, from the chemical reaction models [79] to
physical therapy [80]. For example, Pulsed Electromagnetic Fields are utilized to heal fracture
36
non-unions and treat some bone-related diseases [80, 81], although the specific molecular
mechanisms are not fully understood. However, in this dissertation, I will focus on investigating
the magnetic influence on cell biology.
Literature review shows that both static [29-31, 82] and time-varying [32, 83-85] magnetic
fields can influence tissue culture at the cellular level. Miyakoshi [86] concluded that although a
strong static magnetic field can induce orientation phenomena in cell culture, a static magnetic
field alone does not have significant effects on the basic properties of cell growth and survival,
regardless of the magnetic density. However, with the development of magnetic force-based
tissue engineering, researchers recently have drawn a growing attention to the influence of the
electromagnetic field (EMF), especially to the extremely low frequency electromagnetic field
(<300Hz) because the frequencies of most biological processes are under the range of extremely
low frequencies [85]. Scientists found that compared to a static field, a small pulsed radio
frequency field is more effective in the Ca2+/CaM -dependent myosin phosphorylation system
[86, 87]. Nevertheless, the studies of EMF are mainly focused on connective tissues, bone and
cartilage tissues. For example, it has been proven that EMF (75 Hz, 2,3 mT) can promote
anabolic activites and proteoglycan synthesis in bovine articular cartilage explants [84].
Furthermore, Fassina et al. [83] determined that electromagnetic stimulation with 2 mT flux
density and 75 Hz frequency accelerated SAOS-2 cells’ proliferation and increased calcium
deposition. Komazaki and Takano [88] examined the influence of a low frequency EMF (50 Hz,
5-30 mT), and found that EMF specifically increased the [Ca2+]i of gastrula cells, therefore,
accelerating the rate of morphogenetic cell movements during gastrulation.
These facts motivate us to explore the effects of time-varying electromagnetic field
generated by the stretch device. The magnetic field generated by the apparatus should be weak
37
and at a low frequency (<300Hz) according to the above reviews. But the distribution and
magnitude of the field is unknown, and the effect on cell biology is unclear. The following
sections will illustrate the magnitude and distribution of the flux density B generated by the
electromagnetic field, the factors affecting the field, a potential side effect of the field, and the
cellular response to the field.
3.2 Modeling the Time-varying Electromagnetic Field
In this section, the time-varying electromagnetic field generated by the stretch apparatus is
modelled. The geometrical model is based on a horseshoe-shaped electromagnet MCI-3404
(Mci Limited Co., ON, CA) with the external dimensions 50×48×26mm, which can provide a
time-variant electromagnetic field when excited by a harmonic current. We are interested in not
only the distribution and magnitude of the magnetic flux density, but also the relationship
between the exciting current I and the magnetic flux density B. Theoretically, the exciting
current I determines the distribution and magnitude of density B. On the other hand, density B
controls the magnitude of the magnetic force F. The model is studied in two dimensions,
especially along the horizontal (X) direction because the horizontal magnetic force is served as
longitudinal stretch force, while Y direction is the perpendicular direction on the surface of the
scaffold. Instead of implementing home developed code, which is often time consuming and
requires considerable resources, a commercial finite-element package, COMSOL Multiphysics
3.3 [89], was utilized. All the simulations were performed with the package.
3.2.1 Electromagnetic Theory
Maxwell’s equations can be written as
38
0
B
D
t
BE
t
DJH
(3.1)
The first equation is also named as Ampere’s Circuital Law:
t
DJH
(3.2)
Equation (3.2) states that the curl of the magnetic field intensity H
is equal to the sum of
the current density J
due to the flow of charges and the time derivative of the electric flux
density D
[90]; related variables are shown in Table 3.1.
Table 3.1 Nomenclature
H
magnetic field intensity E
electric field intensity
B
magnetic flux density D
electric flux density
A
magnetic vector potential V electric scalar potential
M
magnetization vector P
electric polarization vector
F
magnetic force J
volume current density
q free charge eJ
external volume current density
permeability permittivity
conductivity angular frequency
d thickness in z direction density
T stress tensor p air pressure
39
In this study case, the electromagnet is supposed to work under low frequency, less than
50Hz. The wavelength for a 50Hz signal is 6000km. On the other hand, the structure size of the
magnets is around 50x50mm, which means the length of the geometry is only a very small
fraction of the exciting wavelength. In this slow variation case, the electromagnetic field is
assumed to be a quasi-static field with 0 tD
. Therefore, Ampere’s Circuital Law (3.2)
simplifies to:
JH
(3.3)
Considering the constitutive relation eJEJ
and the Lorentz force equation
BvEqF
, equation (3.3) for quasi-static electromagnetic field can be extended to:
eJBvEH
)( (3.4)
Combining the constitutive relation, HMHB r 00 )(
, the definition of the
magnetic potential AB
, and electrical potential t
AVE
, Maxwell-Ampere’s law
can be rewritten as:
eJVAvMAt
A
)()( 1
0 (3.5)
Since the exciting current is a harmonic signal, this study case is a time-harmonic case. In
the time-harmonic case, Ampere’s law includes a displacement current Dj , with the
constitutive relation PED 0 , equation (3.5) can be rewritten as following:
PjJVjAvMAAj e
)()()()( 01
002
(3.6)
40
3.2.2 Electromagnetic Model Analysis
Since our objective is to study the field distribution and the flux density B of a quasi-
static magnetic field, we can ignore the coupling between the electric and magnetic field, and
only consider the induction currents that are relevant to the magnetic potential. Moreover, as we
are mainly interested in the B distribution on a horizontal surface (x-y direction), we can study
the magnetic field in two dimensions to simplify the modeling process. In the two-dimensional
case, with only the consideration of induction current, equation (3.6) can be simplified as:
ezzzr JLVAvA
/)()( 110 (3.7)
Equation (3.7) is utilized in the electromagnet subdomain analysis [89]. To perform the
simulation, we make the following assumptions:
The current source is a time-harmonic sinusoidal current with magnitude of 0.4A and
frequency of 1Hz.
The entire coil is considered as a whole block with a constant external current density
corresponding to the current in each single wire. This assumption is more efficient than
modeling each turn of the coil as a separate wire.
The eddy current between the turns in the coil is negligible.
The material of the core is a solid soft iron, ignoring the laminations among the sheets
of the iron core
Based on the above assumptions, we choose the 2D, time-harmonic, quasi-static,
perpendicular induction current, and vector potential application mode in the COMSOL AC/DC
module. The complete set of geometrical and electrical parameters used for the simulation is
listed in Table 3.2 and 3.3. The meshed geometry of the electromagnet model is shown in Fig.
41
3.1. In the graph, there are three subdomains: air, coil (copper) and core (soft iron). The bobbins
between the coil and the core are ignored since they do not significantly influence the magnetic
field analysis. The air space surrounding the feet of the magnet is 30mm, just half the size of a
60mm standard culture dish. Therefore, the graph represents half of a stretch sample; the other
side of the stretch sample is symmetrical. The geometrical dimensions and material properties
of the magnet were provided by the product company. The mesh generator partitions the
subdomains into triangular elements. The finite element model has 2716 elements and the
number of degrees of freedom is 5491. The size and density of the elements are allocated based
on the geometrical dimensions. The areas with more boundaries have smaller elements with
higher density, while the outside areas have larger elements with lower density. This allocation
can reduce the computational expenses and facilitate the calculation procedure.
Table 3.2 Geometrical Parameters
Parameters Expression Description
r_coil 0.2mm Radius of the coil wire
N 550 Turns of the winding
Ww 4mm Width of the winding
Lw 28mm Length of the winding
Wc 12.5mm Width of the core
Lc 38mm Length of the core
w 1mm Width between the core and winding
42
Figure 3.1 Meshed Electromagnet Model
(a)
43
(b)
(c)
Figure 3.2 Distribution of Magnetic Flux Density
(a) Bx component, (b) By component, (c) Bn component
44
Table 3.3 Electrical Parameters
Parameters Expression Description
I_coil 0.4A Current of the coil
J_coil I_coil*N/(Ww*Lw) External current density
5.99e7S/m Conductivity of the copper wire
As the exciting source is a time-harmonic current, the magnetic field distribution of the
study space varies with time. The magnetic poles of the electromagnet alternate between the
two feet during a period. Fig. 3.2 (a), (b), and (c) exhibits the x component, y component and
normal distribution of the maximum flux density in a period at phase zero, respectively. The
units for x and y coordinates are meter. Since the magnetic force along the x direction serves as
the stretch force, the x component of flux density, denoted as Bx, is the main study object. In the
Fig. 3.2 (a), the red arrows act as magnetic field lines representing the magnetic field
distribution. The magnetic south pole (orange) of the horseshoe electromagnet is at the lower
foot and the magnetic north pole (blue) is at the upper foot. The distribution shows that the
magnetic flux density Bx has greater magnitude in the soft iron core than in the air space. The
maximum Bx in the soft iron is 0.204T at the inner corners. However the amplitude of Bx
dramatically reduces to 0.041T at the boundary of the iron core and air space. Except for the
two feet areas, the flux density B in most air space is near zero (green). For example, the
magnitude of the Bx is only 0.015T at 5mm distance of the edge, 0.012T at 10mm distance, and
0.009T at 15mm distance.
45
Figure 3.3 Magnetic Flux Density Bx along y Direction
Since the magnitude of the Bx determines the magnitude of the magnetic force along the x
direction, the effective area of the non-contact magnetic force is limited. In order to apply non-
contact magnetic force, the magnetic rods should be placed within 10mm from the edge of the
electromagnet. Fig. 3.3 shows the distribution of Bx along the y direction (16mm long) at 5mm
distance from the edge of the magnet feet. The plot demonstrates that the magnitude of Bx at
that location is very small, with a maximum value of 0.015T. However, It also shows that the
absolute value of the density Bx is almost symmetrical along the y direction symmetrical axial
at y=0.151, which means if a 16mm long magnetisable rod is put into the field along y
direction, the two sides of the rod will experience symmetric force. This plot will be compared
with Fig. 4.9 in Chapter 4 section 4.6.3. Besides being affected by the flux density, the
magnitude of the force is also determined by the distance between the magnet and the
magnetisable material and the material’s property, which will be discussed in Chapter 4.
46
Although we focus our investigation on the distribution of magnetic flux in the x direction,
we still need to study the normal magnetic flux density B because it is the closest parameter
(2D) to the results of the physical test (3D). For the 2D model, the normal magnetic flux density
B was defined as the total magnitude of the X component and the Y component. The distribution
of the normal magnetic flux density is shown in Fig. 3.2 (c). The simulation results demonstrate
that the normal magnetic flux density B is 0.082T directly beside of the horseshoe magnet. This
result will be compared with the experimental results in the next section.
3.3 Experimental Verification of the COMSOL Model
Serial physical experiments were implemented to verify the COMSOL model of the
horseshoe-shaped electromagnet. In the experiments, the magnet was excited by a pulsed 0.4A
current train with 1Hz frequency, and 50% duty cycle. A Gauss meter with integral Hall Effect
Probe (Model 2010, Magnetic Instruments Inc.) was utilized to test the magnetic flux density B
[91]. The sample time for each test point is one minute. The verification and simulation results
are compared from the aspects of distance, exciting frequency and current.
3.3.1 Flux Density vs. Distance
The first experiment is conducted to verify the relationship between the flux density B and
distance. The corresponding density B was recorded along the x direction every 5mm from each
magnet’s foot. Fig. 3.4 (a) and (b) represent the density B along x direction, at y=0.0143 (south
pole) and y=0.0159 (north pole), respectively. The red curves in the plots are the simulation
results of Bnorm from the COMSOL model. The blue curves are the testing results of flux density
B, which were obtained from the average of twenty peak magnitudes during the sample time.
47
The y axis of Fig. 3.4 shows the absolute magnitude of B for at the south pole (a) and north pole
(b); while the x axis shows the distance from the test points to the edges of the electromagnet.
The graphs demonstrate that the south pole and the north pole possess the same trends.
The trend of B in these plots shows good correlation between the test and simulation results.
Moreover, the field distribution coincides with the Bio-Savart law. According to the Bio-Savart
law, the magnetic flux density produced at a point from an element of length ld
of a
filamentary wire carrying a steady current I is expressed as:
20
4 R
alIdBd R
(3.8)
where is an element of length in the direction of the current, is the unit vector
pointing from to the point, and R is the distance between the point and the current element
. The equation shows that flux density B is proportional to the current I and the reciprocal of
the distance square [90]. However, most of the testing results are, on average, higher than the
simulated results in the ambient area (distance<0.015m) of the magnet except for the zero
position, and lower that the simulated results in the remote area. This phenomenon will be
discussed in the next section.
ld
ld
Ra
ld
48
Simulation vs Testing Results(South Polar y=0.159)
0
0.01
0.02
0.03
0.04
0.05
0.06
0.07
0.08
0.09
0 0.005 0.01 0.015 0.02 0.025 0.03
Distance [m]
Magnetic Flux Density [T]
simulation
testing
(a)
Simulation vs Testing Results(North Polar y=0.143)
0
0.01
0.02
0.03
0.04
0.05
0.06
0.07
0.08
0.09
0 0.005 0.01 0.015 0.02 0.025 0.03
Distance [m]
Magnetic Flux Density [T]
simulation
testing
(b)
Figure 3.4 Simulation Results vs. Testing Results of Magnetic Flux Density
(a) South Pole at y=0.0159, (b) North Pole at y=0.0143
49
3.3.2 Flux Density vs. Exciting Frequency
In order to testify the relationship between flux density and the exciting frequency, the
probe of the Gauss meter was fixed at the lower corner of the north pole (around y=0.0159)
0.005m away from the magnet. The flux density was recorded as the average of twenty peak
magnitude during the sample time, with pulse train frequencies from 1Hz to 5Hz. The test
results show that the flux density ranges from 372.0G to 386.5G, with an average of 375.8G,
and a standard deviation of 5.63 (Table 3.4). The results demonstrate that frequency variations
in the lower range (1~5Hz) have little effect on the magnitude of flux density, which coincides
with the simulation assumption that the eddy current between the wires in the coil is negligible.
Due to the assumption of eddy current, the simulation results show that exciting frequency does
not affect the flux density.
Table3.4 Magnetic Flux Density vs. Exciting Frequency
Frequency
[Hz] 1 2 3 4 5
Flux Density
[G] 386.5 374.7 374.5 372.0 377.0
STD DEV 5.63
50
3.3.3 Flux Density vs. Exciting Current
The relationship between the magnitude of the exciting current and the flux density B was
also investigated in this section. In the interface of the apparatus (Fig. 2.10), only the stretch
related parameters are presented to the users, such as stretch frequency, duty cycle, and strength.
However, all of these mechanical parameters are fulfilled by defining related electrical
parameters. For example, the magnitude of stretch strength is directly controlled by exciting
current. When the stretch strength is set as 95%, the magnitude of the exciting current is 0.4A,
while when the stretch strength is set as 50%, the current is 0.2A.
Exciting Current vs Flux Density
0
50
100
150
200
250
300
350
400
10 20 30 40 50 60 70 80 90
Exciting Current (%)
Magn
etic
Flux
De
nsity
[G]
Figure 3.5 Exciting Current vs. Flux Density
In order to study the relationship between the exciting current and the flux density, the probe
of the Gauss meter was fixed at the lower corner of the north pole at the distance of 0.005m
from the magnet. The flux density B was recorded when the stretch strength increases from10%
51
to 90% by every 10%. Fig. 3.5 shows the test results. It illustrates that the flux density decreases
linearly as the strength of the exciting current decreases. The standard deviation for each test
point is very small, ranging from 0.09 to 1.40. However, the simulation results demonstrate that
flux density is proportional to the exciting current. The test results agree with the simulation
results that the flux density B is proportional to the exciting current I in the low frequency
range. Therefore, the test results verified the design principle that the magnetic flux density and
force can be controlled by controlling the exciting current.
3.3.4 Comparisons of Simulation and Experiment Results
A series of physical experiments are implemented to test the relationships between the
flux density B and the distance, the exciting current and frequency. Although the experiment
results agree with the simulation results, there are still some differences. These differences will
be discussed in the following:
Waveforms
For testing, the exciting current is a pulse train with 1Hz frequency, 50% duty cycle, and
0.4A amplitude. For simulation, we used a time-harmonic sinusoidal current with the same
frequency, duty cycle, and amplitude, but with different shapes. In experiments, the magnetic
field generates when the pulse train switches from OFF time to ON time, and disappears when
the pulse train switches from ON time to OFF time, but the magnetic polarities of the
electromagnet stay the same during the electrifying time. However, for modeling, as the
exciting current is a time-harmonic sinusoidal signal, the magnetic field exists in the testing
space during the whole period, while the magnetic polarities of the magnet alternate during
52
every half period. Although the waveforms of the simulation and the testing are different, the
maximum amplitudes of the two waveforms are the same. Therefore, the simulation results
shown in Fig. 3.2 (the maximum flux density in a period) can represent the real distribution.
Dimensions
When we take a close look at Fig 3.3 (a) and (b), we can see although the testing results
and the simulation results share the same trend, there are three clear differences: (1) the testing
result is lower than the simulation result immediately adjacent to the magnet; (2) except for that
first point, the testing results are slightly higher than the simulation results close to the magnet
(distance<0.015m); (3) the testing results are slightly lower than the simulation results further
away from the magnet (0.015m<distance<0.03m).
The first difference may be caused by the thickness of the test probe. Since the probe of
the Gauss meter has a thickness, the first testing result is not the real density B at zero distance
(the boundary of between the core and the air). There is a small gap between the electromagnet
and the test probe. This may be the reason why the first testing result is lower that the
simulation result as density B decreases dramatically outside of the core. For the second
difference, the simulation results are obtained from a two-dimensional model, while the
experiment results are based on a real three-dimensional magnet. When the probe is positioned
at a specific place, it senses the flux density B from three dimensions (including Bz, density
along z direction); however, the simulation model did not consider this parameter. Therefore,
most of the testing results are slightly higher than the simulation results in the ambient area of
the magnet. With the increase of distance, density B radiates toward different directions, and the
difference between the simulation and testing gets smaller. For the third difference, when the
53
distance from the magnet is greater than 0.02m, the flux density is much lower than 0.01T. In
this situation, the sensitivity of the probe and the test errors greatly influence the experimental
results.
Despite the above differences, the testing results and the simulation results still agree
with each other well. Therefore, the modeling provides a useful tool to optimize the density
distribution and to calculate the non-contact magnetic force.
Figure 3.6 Magnetic Flux Density Distribution of the Stretch Device
(1-Electromagnet, 2-Shadow Area, 3-Scaffold, 4-Steel Bar)
The simulation results show that the spatial distribution of density B is non-uniform. The
ambient area of the magnet (distance<0.01m) has stronger density than other space, shown in
Fig. 3.4. We can apply this distribution feature to the designed apparatus. When a culture dish is
positioned between a pair of the modeled magnets (1), the shadowed areas in Fig. 3.6 (2) have
stronger density. If a pair of magnetic bars (4) is secured in these areas and mounted at the two
ends of an engineered tissue patch, the patch can experience a controllable stretch. Meanwhile,
54
the tissue patch (3) is exposed to a very weak magnetic field. The density B is less than 0.02T,
over a distance of 10mm (the edge of the shadow area). The uniaxial stretch apparatus described
in Chapter 2 is designed based on the above theory. The stretching force can be calculated
directly by the modeling, since the flux density has been verified by the physical experiment.
Hence, the time-varying magnetic field offers a promising approach to apply controlled
mechanical conditions to the tissue culture process via the non-contact magnetic force.
3.4 Effect on Cell Biology
The above analysis shows that magnetic field distribution along the x direction is non-
uniform. The maximum magnitude of the density B is about 0.08T, while the density reduces to
only 0.04T near where the steel bar is secured. The density B in most of the tissue culture area
(Fig. 3.6 (3)) is very low (B<0.01T). Two bio-experiments were performed to test the influence
of the magnetic field on cell biology. One experiment tests the cell proliferation and the other
investigates the cell morphology.
3.4.1 Cell Proliferation
Cells and Cell Culture
To test the magnetic influence on the cell proliferation of engineered tissue patches, we
cultured rat smooth muscle cells on a gelatine sponge scaffold. The Gelfoam was fully soaked
in culture medium and incubated for three days before cell seeding. The mounting trays were
gas sterilized and placed in 60mm Petri dishes. The soaked scaffold (without clamps) was then
secured in the center groove of the mounting tray. Cell suspension (6×106 cells/250μl culture
55
medium) was then seeded into each Gelfoam scaffold (40x20x10 mm). The cell seeding
procedure followed previously established methods [6], and the culture medium was gradually
added until the scaffold submerged.
The assembled dishes were placed at 37℃ in a humidified incubator with 5% CO2 and
95% air for twenty-four hours for cell adhesion. Three samples were subsequently placed on the
newly designed bio-stretch apparatus with the modeled magnets. Since no stainless steel bars
were attached at the ends of the scaffold, the engineered tissue patches were only affected by
the time-varying magnetic field. The field was intermittently applied to the patches, alternating
between 15minutes of stretching time and 30minutes of rest time. The exposure time (15
minutes) referred previous studies [92, 93]. The frequency and strength of the exciting current
were set to be 1 Hz and 0.4A, respectively. In order to obtain a stronger magnetic field, the
strength of the current was set very high. However, the control groups were placed in the same
incubator without being exposed to the time-varying magnetic field. The culture medium was
changed every other day.
DNA Assay
To determine the number of cells on each engineered tissue patch, the total DNA was
extracted using the DNeasy Blood&Tissue kit (Qiagen) following the manufacturer’s
instructions on days 3 and 14 after exposure to the magnetic field. The cell-seeded gelatine
sponge was lysed using the lysis buffer and the lysate was loaded onto the DNeasy mini spin
column. Total DNA was then eluted into TE buffer and read at 260 nm. Table 3.5 compares the
average and standard deviation of DNA at a magnetized condition with the control group from
the two time points (Day 3 and Day 14).
56
Results
Table 3.5 shows that the control group had a slightly higher number of cells than the
magnetized samples. However, the T-test results for two groups show that there is no significant
difference among these groups because P>>0.05. The statistic power analysis was also
performed by SigmaStat software. The calculation results show that the powers of these two
groups are 1.0 and 0.99, respectively, which means there is no significant difference between
the magnetized group and the control group. Besides the analysis results, condensation was
observed on the lids of the magnetized group, which suggests that the temperature of the
magnetized group was different from that of the control group.
Table 3.5 Cell Numbers of Magnetized and Control Samples
3 Days 14 Days
Magnetized
(n=3)
Control
(n=3)
Magnetized
(n=3)
Control
(n=3)
DNA Ave
(μg/ml) 56.2 68.9 69.7 72.7
DNA Std 1.36 0.93 0.33 0.57
T-Test P=0.25 P=0.47
57
3.4.2 Cell Morphology
Cells and Cell Culture
To evaluate the magnetic influence on cell morphology, we used rat smooth muscle cells
(SMC) and mouse bone marrow stromal cells (BMSC) as cell sources. The cells were seeded
directly to a 35mm culture dish. Cell suspension with 0.2×106 cells was seeded into each dish.
After culturing 24 hours, three samples were placed into the same defined electromagnetic field
as section 3.4.1. The control dishes were placed into the same incubator without the magnetic
field. The lag period was intended to provide the cell enough time to recover from isolation
procedure and enable attachment.
Results
We learn from the modeling that the magnetic field distribution of the electromagnet is
non-uniform. Therefore, we only observed the cells within 10mm to the magnet, where the
magnetic density is greater than 0.01T. Fig. 3.7 and Fig. 3.8 show the cell morphology of SMC
and BMSC, respectively. The images were taken by a NIKON Ti-S microscope for phase
contrast. The photographs are grouped into two columns. In Fig 3.7 columns (a) and (b),
comparison is made on the morphology of SMC after 48 hours and 72 hours; similarly, in Fig
3.8 columns (a) and (b), comparison is made on the morphology of BMSC after 48 hours and 72
hours, respectively. In each column, the control image is on the top panel while the image of the
magnetized sample is on the bottom panel. However, no significant difference in cell
morphology was observed for either cell type at the two time points by the observers who were
blinded to the study methods, which means that the time-varying magnetic field with the density
of 0.08T at zero distance has little effect on cell morphology.
58
(a) (b)
Figure 3.7 Cell Morphology of SMC
(a: 48hrs-SMC, b: 72hrs-SMC, magnification: 200x)
59
(a) (b)
Figure 3.8 Cell Morphology of BMSC
(a: 48hrs-BMSC, b: 72hrs-BMSC, magnification: 200x)
3.5 Heating Effect of the Electromagnet
When performing the bio-experiments, condensation was observed on the lids for the
magnetized group. This drew our attention to a possible side effect of heating when using an
electromagnet, a problem because temperature is an important parameter in tissue culture.
Normally, the temperature remains at 37℃ in the incubator. However, when we simulated the
electromagnet model, the heat dissipation problem was ignored, so that the increase of the
temperature cannot be predicted. Therefore, the variation of the temperature was tested by
experimentation.
3.5.1 Experiment Results
Under Room Temperature
The Hall Effect Probe of the Gauss meter (Model 2010, Magnetic Instruments Inc.),
which we used to verify the flux density B, can also test temperature. When we studied the
60
relationship between flux density B and exciting frequency f, the probe was fixed at 5mm from
the edge of the magnet. The probe not only sensed the variation of the flux density, but also
recorded the variation of the temperature. Fig. 3.9 illustrates the temperature variation after the
electromagnet was excited. The electrify parameters are set as 1Hz frequency, 95% strength,
and 50% duty cycles. The plot shows that the temperature increased 1.3℃ after 9 minutes
exciting, and remained stable.
Temperature vs Time
25
25.5
26
26.5
27
0:00
:00
0:00
:27
0:01
:38
0:02
:06
0:03
:17
0:03
:44
0:05
:08
0:05
:35
0:06
:02
0:07
:08
0:07
:35
0:08
:46
0:09
:13
0:09
:40
Time [h:mm:ss]
Tem
per
atu
re [
℃ ]
Temperature
Figure 3.9 Temperature vs. Exciting Time under Room Temperature
Under Incubator
The above test was performed under room temperature, while the real working
environment for the electromagnet is in the incubator. Therefore, a precision thermometer UNI-
T (UT328, UNI-TREND technology Inc., Guangdong) was utilized to test the temperature
variation under working conditions. The thermometer has two thermocouple input channels.
One thermocouple (t1) was fixed in the middle of a pair of electromagnets (the center of the
61
stretch board) to sense the temperature variation caused by the magnets; another thermocouple
(t2) was fixed at the metal shelf in the middle of the incubator to sense the temperature of the
incubator. After the test points fixed, the stretch board was electrified. The stretch plan was set
the same as in the bio-experiments in section 3.4 (electrifying 15 minutes and resting 30
minutes, 1Hz, 95% strength, 50% ON time). The thermometer was placed in the incubator for
11 hours. After three hours and forty minutes (five periods of the stretch plan) for satiability, it
started to record the two test points’ temperature every minute. Fig. 3.10 records seven and half
hours (10 successive periods) data.
Fig. 3.9 and Fig. 3.10 show clearly that the temperature increases when the
electromagnet is electrified. Both experiments verified that the local temperature around the
magnets is about 1.0℃ higher than the environment. Fig. 3.9 emphasizes the short time
increase trend under room temperature. The temperature increased 1.3℃.
Temperature vs Time
35.5
36
36.5
37
37.5
38
38.5
1:0
7:4
7
1:2
6:4
7
1:4
5:4
7
2:0
4:4
7
2:2
3:4
7
2:4
2:4
7
3:0
1:4
7
3:2
0:4
7
3:3
9:4
7
3:5
8:4
7
4:1
7:4
7
4:3
6:4
7
4:5
5:4
7
5:1
4:4
7
5:3
3:4
7
5:5
2:4
7
6:1
1:4
7
6:3
0:4
7
6:4
9:4
7
7:0
8:4
7
7:2
7:4
7
7:4
6:4
7
8:0
5:4
7
8:2
4:4
7
Time [hh:mm:ss]
Tem
pe
ratu
re [
℃ ]
t1
t2
Figure 3.10 Temperature vs. Exciting Time in incubator
62
Fig. 3.10 shows the temperature fluctuation in the incubator. Compared with the baseline
t2, t1 shows more severe fluctuation. During this test period, the temperature for t1 ranges from
37℃ to 37.9℃, with an average of 37.4℃; while the temperature for t2 ranges from 36.℃
to 36.7℃, with an average of 36.4℃. The temperature influence on cell biology was
investigated in the following section.
3.5.2 Bio-experiment Verification
As the experiment results show that the electromagnets causes the local temperature
fluctuated during electrifying, while temperature is an important culture parameter for cell
growth. The biological influence of temperature needs to be clarified.
Cells and Cell Culture
To test the influence of temperature on the cell proliferation of engineered tissue
patches, we again used the gelatine sponge (Gelfoam: Pharmacia & Upjohn Co., Kalamazoo,
MI) as the scaffold and rat smooth muscle cells as the cell source. The experiment followed the
same cell-seeding protocol described in section 3.4.1. However, we used only half the number
of the cells. As the experiment was performed in 35 mm Petri dishes, the size of the scaffold
was only a half of the experiment in section 3.4.1. Hence, the cell suspension (3×106
cells/125μl culture medium) was seeded into each Gelfoam scaffold (20x20x10 mm). The cell-
seeded patches were then assembled into 35 mm culture dishes. All the samples were first
placed into a humidified incubator with 5% CO2 and 95% air at 37℃ for twenty-four hours.
63
The test and control groups were then placed into two different humidified incubators with 5%
CO2 and 95% air for three days. The temperature for the test group was 38℃, while the control
group’s temperature was 37℃.The number of samples in each group was three.
DNA Assay
A DNA assay was again used to determine the number of cells on each engineered tissue
patch, following the same procedure as in section 3.4.1. Table 3.6 compares the average and
standard deviation of DNA at a high temperature condition (38℃) with the control group
(37℃).
Table 3.6 Cell Numbers at 38℃ and at 37℃
High Temperature 38℃
(n=3)
Control 37℃
(n=3)
DNA Ave
(μg/ml)
48.2 52.0
DNA Std 2.30 5.26
T-Test P=0.32
Results
Table 3.6 shows that the total cell number of the control group is slightly higher than the
group cultured under higher temperature. However, the T-test result for two groups shows that
there is no significant difference among these groups because P>>0.05. This result is in
64
accordance with the result obtained from section 3.4.1. The phenomenon may explain that one
degree of temperature difference does not significantly affect cell proliferation. In section 3.4.1,
the magnetized group was not only exposed to the time-varying magnetic field, but also affected
by the temperature variation. However, these two parameters also do not significantly affect the
experiment results.
3.5.3 Effect Factors
When an external electric current passes through an electromagnet, the heat conductivity
is known to be coupled with the electromagnetic field simulation. Therefore, the temperature
variation was examined.
There are several physical parameters affecting the heating problem of a pulsed
electromagnet, such as permeability µ, conductivity , and permittivity Ɛ. Heat is generated not
only in the coils but also from the core [65-68]. For example, when the iron core is exposed to a
pulsed magnetic field, the permeability µ of the material will change. The temperature
dependent parameter will introduce a phase shift between the magnetic flux density B and the
magnetic field strength H that causes the hysteresis pneumonia. Joule Heating can be called
resistance heating that is mainly heat loss of the coils. It is determined by the local current flux
and temperature-dependent conductivity. In a Joule Heating study, the heat transfer problem
couples with the electronic current problem. The calculation of the heat source is based on the
equation 2
VQ . However, the thermal energy in turn changes the electrical conductivity
of the coil. The temperature-dependent conductivity can be expressed as
))(1(1
00 tt , where 0 ( m ) is a reference resistivity at a reference temperature
65
t0(K), and α is proportionality constant (K-1) for the temperature dependence. The temperature-
dependent conductivity causes the variation of the exciting voltage, which results in the
coupling between the temperature and electromagnetic fields.
In future modelling, a multiple physical aspects need to be studied. The model should
consider not only magnetic field, but also heat transfer. The above effect factors should be
considered into the modelling, which is useful to accurately describe the real phenomena. To
reduce the effect of temperature fluctuation in later experiments, there are two methods can take
into account. First, reduce the temperature of the incubator 0.5℃ lower than the previous set
number, which can effectively reduce the average temperature of the stretch group, but the
control group will also be affected. Another way is to place the electromagnets onto a metal
base such as aluminum, which can disparate the heating and not affected by the magnetic field.
Therefore, the local temperature of the electromagnet can lower down.
3.6 Conclusions
The time-varying electromagnetic field generated by the stretch device is quantitatively
studied in this chapter, both through simulation and experimentation. The magnetic field is
modeled by COMSOL multiphysics 3.3 and verified by the Gauss meter Model 2010. The
investigation shows that the distribution of magnetic flux density along the stretch direction (X
direction) Bx is non-uniform; the magnitude of the density is proportional to the exciting current
I; and the magnitude of the density is not related to the exciting frequency at low frequency
range (1Hz to 5Hz). The relevant biological experiments demonstrate that the weak, low
frequency magnetic field generated by the stretch device does not significantly affect cell
proliferation and cell morphology. However, the difference between the exam group and the
66
control group is caused by the influence of both magnetic field and temperature field as the
electromagnet has heat dissipation problem.
67
Chapter 4
Effect of the Uniaxial Cyclic Stretch on
Cell Biology
After constructing the biostretch apparatus (Chapter 2) and studying its magnetic influence
(Chapter 3), we will investigate the mechanical stimulation provided by the device in this
chapter. We will quantitatively study the influence of the mechanical stimulation generated by
the developed apparatus from the aspects of strain distribution, strain rate, and stretch force.
Relevant bio-experiments are conducted to validate the performances of the apparatus.
This chapter is organized as follows. A brief introduction of the current state of stretch
applications is presented in section 4.1. Section 4.2 validates the apparatus’ performance by
examining cell proliferation and cell orientation. Section 4.3 investigates the strain distribution
on the surface of the scaffold with a nondestructive method. Section 4.4 studies the strain rate of
the apparatus with Photron’s high-speed camera Ultima APX. Section 4.5 calculates the stretch
force provided by the apparatus. Section 4.6 draws conclusions from the above research.
68
4.1 Introduction
Tissue engineering shows great potential for treating diseases by providing functional cell-
based substitutes for human tissues [6, 94, 95]. In addition, engineered tissues can also serve as
in vitro tissue equivalents for drug testing and development, and even for studying the tissue
functions that can expand our understanding of tissue biology [69, 96, 97]. However, in
conventional three-dimensional tissue culture systems, engineered tissues could not develop the
similar structure as native tissues. Compared with engineered tissues, native tissues live in a
much more complicated environment in vivo. For example, connective tissues such as tendon
and ligaments experience tensile loading, while native cardiac tissue contracts under the control
of electrical pacing signals. The specific constructs of tissues are built to adapt their
environment. Currently, most engineered tissues are cultured under a relevantly static
environment. Their cell density and mechanical properties in tension and compression are
substantially lower than those of native tissues [59]. Therefore, external physical stimuli are
necessary for the culture systems to mimic various aspects of the in vivo environment.
Different types of cells are sensitive and responsive to different mechanical stimuli. As
described in Chapter 1 section 1.2.1, smooth muscle cells usually experience tensile forces [7].
Since the designed apparatus is developed for uniaxial cyclic stretch, smooth muscle cells are
chosen as the cell source for all the relevant bio-experiments. Moreover, the types of the
uniaxial biostretch apparatus are reviewed in Chapter 2 section 2.2. Based on these devices,
related bio-experiments have demonstrated that cyclic stretch is a potent stimulus, which can
improve cell proliferation, stimulate extracellular matrix production and the mechanical
properties of tissues [6, 12, 24, 24, 54, 55, 59, 62, 64, 97, 98].
69
The group of Vandeburgh [12, 54-56] studied the effects of uniaxial stretch for decades.
Their results show that repetitive stretch/relaxation for eight days increased the elasticity, mean
myofiber diameter, and myofiber area percent of human bioartifical muscles (HBAMs). In
addition, they noted that the muscle cells generated increasing internal forces during formation.
Studies also show that uniaxial cyclic stretch improved the structure and function of cultured
avian and rodent muscle [54, 99]. Eschenhagen et al. and Zimmermann et al. [16, 24]
demonstrated that engineered heart tissue (mixture of neonatal rat cardiac myocytes, collagen I
and matrix factors) displayed important hallmarks of differentiated myocardium when subjected
to unidirectional cyclic stretch. Shimko et al. [98] used murine embryonic stem cells (ESC) to
examine the direct effect of mechanical loading (uniaxial stretch) on the differentiation of ESC-
derived cardiomyocytes, and found that mechanical loading significantly affected gene
expression. Akhyari et al. [6] utilized human heart cells to demonstrate that uniaxial stretch
enhanced the formation of a three-dimensional tissue-engineered cardiac graft.
The above influences were investigated by different custom-built devices. Due to the
diversity of the experimental devices, the comparison of all the mechanical parameters provided
by the devices is difficult. The most commonly studied mechanical parameter for these devices
is strain. In our study, we employed noncontact magnetic force as the cyclic uniaxial stretch
force, and provided two different stretch methods. Compared with the previous device [66], we
can precisely controll the mechanical parameter, strain, by confining the distance between the
stoppers in the mounting tray. However, uniaxial stretch can vary with other parameters, such as
stretch pattern, strain distribution and rate, stretch force and frequency, which will be discussed
in the following sections. These parameters will all contribute to cell orientation. Nevertheless,
70
before studying these mechanical parameters, the validation of the apparatus will first be
examined by the bio-experiments.
4.2 Validation of the Uniaxial Cyclic Stretch
Although many results of strain-induced cell proliferation, matrix production, and
increased mechanical properties have been widely reported in previous literature review [13,
58], the performance of the designed apparatus still needs to be validated by the biological
experiments. Since cell proliferation has been utilized to study the influence of the magnetic
field and temperature, the same method will be used to test the influence of uniaxial stretch.
4.2.1 Cell Proliferation
Cells and Cell culture
In the experiment, the gelatine sponge (Gelfoam: Pharmacia & Upjohn Co., Kalamazoo,
MI) was again used as the scaffold and rat smooth muscle cells served as the cell source. All the
stretch related accessories, such as clamps, steel bars, and boards, were sterilized. The
experiment followed the same cell-seeding protocol described previously. The cell suspension
for each Gelfoam scaffold (30x10x10 mm) was 3×106 cells/125μl culture medium. The scaffold
and clamps were assembled in the hood with tweezers. The assembled scaffolds were then
mounted into 35mm Petri dishes, and cells were seeded on the top of the Gelfoam. The samples
were then placed at 37℃ in a humidified incubator with 5% CO2 and 95% air for twenty-four
hours. Five sample dishes were subsequently placed on the newly designed bio-stretch
apparatus, and another five samples were used as controls in the same incubator. The stretching
71
group was exposed to intermittent stretch, alternating between 10 minutes of stretch time and 10
minutes of rest time for three days. The parameters of the mechanical stimulus are defined as
1Hz, 50% strength, 50% ON time and 20% strain. The culture medium was changed every day.
DNA Assay
After three days of stretching, DNA assay was utilized to determine the number of cells
on each engineered tissue patch following the same instruction described before. Table 4.1
compares the average and standard deviation of DNA concentration for each group.
Table 4.1 Cell Numbers of Stretched and Control Samples
Uniaxial Stretched
(n=5)
Control
(n=5)
DNA Average
(μg/ml)
60.0 36.2
DNA Std 5.37 4.99
T-Test 0.0009
Results
Table 4.1 shows that the average of the DNA concentration is much higher in the
stretched group than in the control group. The DNA concentration between the stretched group
and the control group is statistically significant (p<<0.05).The results demonstrate that this
device can provide uniaxial cyclic stretch to the tissue construct, and thus influence cell
proliferation. This result was in accordance with other research groups’ results [6, 55, 97].
72
However, the experiment results are the response to all the characteristics of a uniaxial cyclic
stretch regime (deformation magnitude, stretch strength, stretch frequency, and even stretch
methods). The independent contributions of each mechanical parameter are unclear.
4.2.2 Cell Orientation
After studying the stretching effect on cell proliferation over a period of in four days,
the influence of stretch on cell orientation was studied for a longer term (2-weeks).
Cells and Cell culture
In this experiment, the gelatine sponge (Gelfoam: Pharmacia & Upjohn Co., Kalamazoo, MI)
was again used as the scaffold and rat smooth muscle cells served as the cell source. The
experiment followed the same procedure described in section 4.3.1., with the only difference
being that the culture period lasted 14 days. The sample number for each group is three.
Hematoxylin and Eosin (HE) Staining
Fourteen days after being seeded with cells, the engineered tissue patches (stretched (n=3);
control (n=3)) were carefully removed from the clamps in the culture dishes. The samples were
then fixed in 10% neutral buffered formalin for 24 hours. Using standard methods, the samples
were embedded in paraffin blocks, with care being taken to maintain their final orientation from
the apparatus.
Two paraffin sections were cut out from each sample, parallel to the surface of the samples.
HE staining was used to examine the gross tissue structure. Hematoxylin stains cell nuclei and
polyribosomes, giving nucleic acid a blue color. Eosin stains protein and cytoplasm, giving it a
73
pink color. The stain images were observed under a Nikon Eclipse 80i fluorescence microscope.
Fig. 4.1 compares the stretched and control sample under the magnification of 100x and 200x,
respectively.
(a)
(b)
Figure 4.1 Gross Structure of the Engineered Tissue
(a) stretched (left: 200x, right: 100x ); (b) control (left: 200x, right: 100x )
74
Results
The images for both the stretched group and control group show that fourteen days later
cells still alive on the scaffold. Further, the visual observation demonstrates that the stretched
constructs have more cells than the control group. Three different images were examined for
each section from vary parts of the scaffold. The number of cells in each image (0.2mm2) was
counted in Adobe Photoshop CS (Adobe Systems, San Jose, California). The stretched group
and control group were averaged and compared (18 images each group) in Fig. 4.2. The
average number of cells was 194 for the stretched group, and 81 for the control group. The T-
test result shows p<<0.05, and marked as * in the graph, which means the stretched group has
significant difference compared with control group.
Cell Proliferation
0
50
100
150
200
250
Ce
ll N
um
be
r a
t 0
.2m
m2
stretched
control
*
Figure 4.2 Comparison of Cell Numbers
(* represent the significant difference between two groups)
Moreover, for the control samples, the cells were scattered among the holes of the
scaffold (Fig. 4.1(b)). For the stretch samples, the cells demonstrated inhomogeneous cell
75
distribution. The cell clusters were parallel along the stretch direction, longitudinally oriented
(Fig. 4.1(a) arrow direction). Additionally, compared with the control group, the stretched
samples exhibited a good cell orientation at the free edges and two ends. Complexes of
multicellular aggregates and longitudinally oriented cell bundles were also found at the ends
and free edge areas (Fig. 4.3).
Figure 4.3 Edge Structure of the Engineered Tissue
(a) control; (b) stretched (magnification: 100x )
4.2.3 Discussion
The results of these two bio-experiments validated the performances of the apparatus. The
device is capable to provide uniaxial cyclic stretch during tissue culture with well-defined
mechanical parameters (strain amplitude, frequency, strength, and duty cycle), which can affect
cell proliferation and cell orientation. However, the results reflected the influence of all
characteristics of a stretch regime, not a single parameter.
76
Cell Reorganization
The HE images show that without external stimulation, cells clustered among the holes of the
scaffold randomly. However, under cyclic stretch, these tissue-like bundles are not uniformly
distributed along the surface of the scaffold. Cells aggregated at the free edges and ends of the
scaffold, while the centre part had fewer bundles. This phenomenon can be explained as
multiple stimuli existed during stretch. The scaffold was not only subjected to uniaxial stretch,
but also strain-induced fluid flow around and within the porous scaffold. Therefore, this culture
system actually induced not only cyclic stretch but also shear stimulation. Additionally, due to
Poisson’s effects, when the substrate undergoes one dimensional stretch, the other two
dimensions will experience compression. During a stretch/relaxation period, the medium inside
the scaffold may experience a squeeze/inhale period, which in turn improved the mass
transportation condition [17][59]. In future studies, the effect of each mechanical parameter
needs to be isolated to help researchers determine the mechatransduction of the cells.
Variation of Mechanical Property
For the long-term culture, it was observed that the scaffolds maintain a high degree of
elasticity during the first seven days. Later, a stronger force (60%) is needed to obtain the same
strain magnitude, which indicates that the stiffness of the structure changed during culture. This
kind of variation has been observed and recorded by other groups [55, 97].
4.3 Study of Strain Distribution
From the literature review listed in table 2.1, we know that the most frequently studied
parameter of cyclic stretch is strain. The average strain Ɛ provided by the device is defined
77
as 0/ LL . L0 is the distance between two symmetric inner stoppers, shown in Fig. 2.6. For
symmetric stretch, L is the total deformation of the scaffold, which is twice the distance from
the inner stopper to the outer stopper. For asymmetric stretch, L is the deformation of the
movable end of the construct. Unlike an elastic rigid-body, the scaffold is a vesoelastic material.
Although the average strain of the construct is clear, the strain distribution at the surface of the
scaffold is unknown. Therefore, a nondestructive method [70][100, 101] was utilized to
investigate the full-field surface strain distribution of the three-dimensional scaffold.
4.3.1 Measurement Methodology
To study the strain distribution, samples of both Gelfoam (41 x 20 x 7 mm, n=2) and GE
RTV 6166 silicone (38 x 16 x 3 mm, n=11) were used as scaffolds. Compared with Gelfoam,
which is porous and has a heterogeneous property, the GE RTV 6166 silicone is a homogeneous
material, allowing it to serve as a baseline to compare to. The Gelfoam samples were marked
with dots at 2mm intervals using a fine-tipped permanent marker; while the RTV samples were
marked by dark microbeads. These marks separated each RTV silicone sample and Gelfoam
sample into 16 sections along the stretching axes (Fig. 4.4). The rectangular areas in Fig. 4.4 are
the analysis area. After marking, the scaffolds were affixed to stainless steel clamps with
silicone glue. Scaffolds were then soaked in water for at least 24 hours. The final dimensions of
the scaffolds with clamps were slightly different due to the gap made by the glue and shrinkage
when the scaffolds were inserted in water, ranging from 40x18x7 mm to 40.5x18x7 mm for
Gelfoam, and 40x16x3 mm to 41.5x16x3 mm for RTV silicone. The scaffolds were then placed
in the aforementioned apparatus for cyclic stretch. Two different stretch methods were applied:
symmetric stretch and asymmetric stretch, which were discussed in Chapter 2 section 2.3.2.
78
In these trials the outer stopper pins were not used. For asymmetric stretch, the inner stopper
pins on the fixed end were placed to secure the scaffold directly against the side of the culture
dish, and on the free end were placed to prevent the scaffold from deforming to less than its
original length. For symmetric stretch, the inner stopper pins were placed only to prevent the
scaffold from deforming to less than its original length, and the scaffold was allowed to stretch
the full length of the culture dish. In this way, the stretch magnitudes of the two methods were
the same. For Gelfoam, the stretch parameters of the biostretch apparatus were set with
frequency of 1Hz, ON time of 50%, and strength of 35% of the maximum value. For RTV, the
parameters were set with frequency of 1Hz, ON time of 50%, and strength of 60% of the
maximum value. The average strain was about 20%.
Digital photographs of the scaffold in the stretched and unstretched states were taken with a
D50 digital camera (Nikon Corporation, Tokyo, Japan) and a Tamron SP 90mm macro lens
(Tamron Co., Ltd., Saitama, Japan). Photographs had a resolution of 60 pixels per millimetre.
The photographs were edited with Adobe Photoshop CS (Adobe Systems, San Jose, California)
using the High Pass and Threshold tools to identify dots. The centroid of each dot was
determined using Scion Image (Scion Corporation, Frederick, Maryland), and distances
between successive centroids along the stretch axis were calculated to determine the distribution
of strain, defined as deformation (∆L) divided by the original distance between dots (L0). For
each distance along the stretch axis, the strains at several equally spaced locations on the
perpendicular axis were calculated and averaged, and the standard distribution of those strains
was determined.
79
(a) (b)
Figure 4.4 Images of the Gelfoam and GE RTV 6166
(a) up: unstretched RTV, down: stretched RTV
(b) up: unstretched Gelfoam, down: stretched Gelfoam
4.3.2 Results
We investigated two types of strain distributions, global strain and local strain distributions.
Global strain distribution refers to the relationship between the strains of different sections;
local strain refers to only the strain of a specific section.
Global Strain distribution of Gelfoam
Based on the above method, we investigated the global strain distribution in both the stretch
axis (x-axis) and the perpendicular horizontal axis (y-axis). In general, strain distribution was
found to be non-uniform, with a high standard deviation, but exhibiting no significant trends
80
along the stretch axis. Fig. 4.5 (a) shows the average deformation between each adjacent pair of
data points for 10 rows of one Gelfoam sample along the x-axis. No statistically different results
were found for global strain distribution of Gelfoam, which is due to high standard deviation in
measurements.
(a)
(b)
Figure 4.5 Strain Distribution of a Typical Sample of Gelfoam
(a) Up: the x- axes; (b) Down: the y- axes
81
A similar graph for strain distribution in the y-axis shows a clearer trend, with the edges of
the Gelfoam deforming least, since they were securely glued to solid metal clamps (Fig. 4.5
(b)). Fig. 4.5 (a) and (b) demonstrates that the scaffold is not only experienced longitudinal
tensile, but also compressed along the lateral direction owing to Poisson’s effects.
Local Strain Distribution of Gelfoam
Compared with the global strain distribution, the deformation of small sections of Gelfoam
(local strain) was far more significant. Photographs of the Gelfoam indicated that shapes of the
dots were visibly distorted during stretching (Fig. 4.6), such that local strain in small regions
(100µm) ranged from negative strains on the order of -5% to large positive strains on the order
of 100%. One explanation for the phenomena is the non-uniform construction of the Gelfoam
sponge, which caused areas with larger surface cavities to exhibit higher strain than areas with
smaller surface cavities. Additionally, parts of the Gelfoam may have torn under tensile stress
and hence exhibited higher strain. In comparison to the global strain (10-20%) and the size of a
typical cell (10-40µm), these local strains are clearly significant. We note that the material
properties of the Gelfoam may be different when the Gelfoam is seeded with cells; this
difference will be discussed in section 4.6.
Figure 4.6 Distortion of a Marked Dot
Left: before stretching; Right: after stretching
82
The significance of the local strains is demonstrated by comparing Gelfoam to RTV silicone,
which has a visibly more uniform construction. Local strains between dots on Gelfoam spaced
2mm apart exhibit much higher standard deviation than those on RTV silicone, with a standard
deviation of about 30% of the average strain in comparison to less than 10% for RTV silicone.
A visual comparison of typical strain distributions for the two kinds of scaffolds, as in Fig. 4.7,
supports this claim. Since these local strains were calculated by measuring distances between
centroids of dots, they are in effect average values, and the standard deviation for local strains is
actually even higher than calculated.
Figure 4.7 Strain Distribution for the Middle Rows of the Gelfoam and RTV Silicone
Samples
Symmetric and Asymmetric Stretch
Two different stretch methods, asymmetric and symmetric stretch, are compared for two
kinds of samples. However, there are no significant trends along the stretch axis under different
stretch methods. But the symmetric stretch method provides a good chance for real-time strain
monitoring since the sample could stay centered below a camera. Moreover, other simple
83
physical variables, such as displacement and speed of particles, may be easier to correlate with
differences in cell growth. For example, under one-way stretch, the free end is displaced further
and has a higher average speed than the fixed end, which may correlate with cell growth.
Edge Effects
We note that the strain distribution in the edge sections, where the clamps are attached, is
heavily dependent on the method of application of force, and is subject to high measurement
error. Before utilizing the clamps described in Chapter section 2.3.2, other clamps were tested.
When force was applied to metal bars inserted through the Gelfoam, strain measured at the
edges was less than half of that measured at the centre. Conversely, when clamps were affixed
to the Gelfoam, strain at the edges appeared more than double the strain measured at the centre.
These are genuine differences in strain; however, they are not necessarily the cause of
differences in cell expression, as there are many other differences between the centre and edges
of the Gelfoam. We also note that the substrate undergoes deformation in the vertical direction
near the edges, which causes strain to appear different from its actual value. For these reasons,
we omit the outer 2 millimetres in all of our measurements.
Since the analyzed data was only taken when the Gelfoam was stretched either fully or not at
all, no information was obtained about the partially stretched state of Gelfoam, and hence strain
rate distribution is unknown. Previous biomechanics studies show that biological tissues are not
elastic materials, the living tissues are normally recognized as viscoelastic materials with
nonlinear stress-strain characteristics [102]. Instead of the strain magnitude itself, the history of
84
strain affects the stress more. Therefore, the study interest was shifted from strain distribution to
strain rate.
4.4 Study of Strain Rate
From the design principle of the apparatus, we know that when the magnet is electrified by a
pulsed current, the magnet will generate a time-varying magnetic field. When the assembled
dish (Fig. 2.6) is positioned under such a field, a magnetic force will act on the magnetizable
bar, which will then stretch the connected tissue construct longitudinally. When the magnetic
field disappears, the tissue construct will restore to its original shape due to the elasticity of the
scaffold. Under a confined strain value, if the stretch force varies, the strain rate will change.
Therefore, the relationship between stretch force and strain rate at the defined strain magnitude
was investigated.
4.4.1 Measurement Methodology
The engineered tissue was prepared according to the protocol provided in section 4.3. After
the tissue patch was cultured for 16 days, the assembled dish was positioned under the scope of
Photron’s high-speed camera ultima APX (Photron, San Diego, California), and continuously
subjected to symmetric stretch. The stretching parameters were set as: 1Hz, 50% ON time, and
the total deformation of 5mm. As the original length of the Gelfoam sample is 20mm, the
maximum strain is 25%. The stretching process was recorded by the digital imaging system.
The high-speed digital imaging system was set to 500 fps at full mega pixel image resolution
(1024 x 1024). The recording time was 4.1 seconds, with 2048 frames, such that each record
85
includes four successive stretch periods. The recordings were saved in AVI format. Four video
outputs were recorded, with stretch force ranging from 40% to 90%.
4.4.2 Results
The time period of each frame of the video represents 2ms. The images were edited by GIMP
(GNU General Public License). With the help of the calibration picture, the stretch strain rate,
the restore strain rate, and the frequency response of the device for a specific material and strain
were determined from the video.
Stretch Strain Rate
The data show that the stretch strain rate varies under different stretch strengths. Fig. 4.8
(a) and (b) show the stretch rates with different stretch strengths of 90% and 40%, respectively,
and the stretch processes demonstrate very good repetition for the four successive periods.
When the strength of the stretch force reaches 90%, 16ms is needed to reach the total
deformation. When the strength is only 40%, 22ms is needed to reach the same deformation.
Despite the difference in stretch time, the two plots share the trend that the stretch rates vary
and increase nonlinearly during the stretch process. These plots reflect the nonlinear viscoelastic
property of the material or indicate that the stretch force is not constant during the stretch
process. Thus, this stretch force will be calculated in section 4.6.
86
stretch rate at 90% strength
0
0.5
1
1.5
2
2.5
1 2 3 4 5 6 7
frames
de
form
ati
on
[m
m]
round1
round2
round3
round4
(a)
stretch rate at 40% strength
0
0.5
1
1.5
2
2.5
1 2 3 4 5 6 7 8 9 10
frames
def
orm
atio
n [
mm
]
round1
round2
round3
round4
(b)
Figure 4.8 Stretch Strain Rate at Different Stretch Strength
(a) 90% strength; (b) 40% strength
87
The strain rate during the stretch process shares the exactly same plots as stretch rate (Fig.
4.8), since the strain is obtained by deformation (L) divided original length (L0). In the test
range, the maximum strain rate for 90% stretch strength is about 107s-1, while the maximum
strain rate for 40% stretch strength is about 102s-1.
Retreat Rate
When the external force (magnetic force) is removed, the construct will retreat back to its
original place due to the elasticity of the material. The analysis data show that the retreat
process takes 48ms, which is much longer than the stretch process. However, these two plots
demonstrate the same trend, which shows that the retreat process is not affected by what’s
happened in the past but the material’s own material properties. The plots demonstrate the creep
property (time dependent strain rate) of the material during the unloading process. These results
are in accordance with the facts that the scaffold is a viscoelestic material. But the plots (Fig.
4.9) do not match the creep functions of the ordinary linear viscoelastic mechanical models (the
Maxwell model, the Voigt model, and the Kelvin model) [102]. Therefore, the material property
of the engineered construct needs further study.
The scaffold of the sample is Gelfoam, an absorbable gelatin sponge. The material
properties of the Gelfoam may change when it is used as a scaffold, due to biodegradation and
cell reorganization. This non-invasive test method provides a new avenue to monitor the change
of the mechanical properties of the engineered tissue during culture process.
88
retreat rate at 90% strenght
0
0.5
1
1.5
2
2.5
3
1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24
frames
def
orm
atio
n [
mm
]
round1
round2
round3
round4
(a)
stretch rate at 40% strength
0
0.5
1
1.5
2
2.5
3
1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24
frames
def
orm
atio
n [
mm
]
round1
round2
round3
round4
(b)
Figure 4.9 Retreat Strain Rate at Different Stretch Strength
(a) 90% strength; (b) 40% strength
89
Frequency Response
The above data provides us a method to study the frequency response of the device for a
specific material and strain. The stretch force is generated from the pulsed exciting current,
shown in Fig. 4.10 (a). Theoretically, the frequency of the current pulse can reach very high
values, as the maximum source frequency of the timer is 20MHz. However, according to the
recorded data, the stretch time and the retreat time are 16ms and 48ms, respectively. For a 10Hz
pulse train, the response stretch pulse is shown in Fig. 4.10 (b). The shortest response time that
the construct needs for a period is 64ms (Fig. 4.10 (c)). Therefore, the frequency response of the
scaffold is about 15Hz for 25% strain. The data demonstrate that frequency response is related
to not only the stretch force, but also the strain (deformation) and the material properties.
However, this calculation ignores the delay between the exciting current and the generated
magnetic field.
Figure 4.10 Stretch Frequency
90
4.4.3 Discussion
All the collected data were recorded under the real culture circumstance. During the process,
the Gelfoam were soaked in medium. The video shows that during stretching, the fluctuation of
the medium distorted the image, especially at the moment that the metal bar hit the wall of the
culture dish or the stoppers of the mounting tray, which is the major measurement error during
the test. Therefore, it was hard to capture the accurate position of the steel bar at the hitting
momentary. Due to this reason, the deformation of the last frame is not plotted in Fig. 4.8.
This test demonstrated that for a specific material the stretch rate is related to the strength of
the stretch force. However, the stretch rate is not proportional to the stretch force. When the
force doubled (from 40% to 90%), the full stretch time did not reduce to half (from 22ms to
16ms). Additionally, the material property affects both the stretch rate and the retreat rate. In
order to obtain the same strain magnitude, different scaffolds need different initial force to
stretch. This claim has been supported in section 4.4.1. To test the strain distributions of
Gelfoam and RTV silicone, different stretch force were applied. Therefore, this nondestructive
method also provides researchers a tool to real time monitor the variation of the material
property during cell orientation.
4.5 Study of Stretch Force
From section 4.5, we noticed that the stretch force may not be a constant force. In this
section, we will estimate the distribution and magnitude of the magnetic force generated
between a magnetisable bar and a horseshoe-shaped electromagnet. The geometrical model of
the magnet is based on MCI-3404 (Mci Limited Co., ON, CA) utilized in the biostretch
91
apparatus, the magnetizable bar (16x 3.5x 1mm). We focus our model study on two dimensions,
particularly along horizontal (x) direction because the horizontal magnetic force is used as the
uniaxial stretch force. In this model, the two dimensions of the bar are 16 x 1 mm, and s the
distance between the bar and the magnet is 5mm. The simulations were performed using
COMSOL Multiphysics 3.3.
4.5.1 Electromagnetic Force Calculation
In COMSOL Multiphysics, Maxwell’s stress tensor T is used to calculate the magnetic
force act on the surface of the object. Cauchy’s equation of continuum mechanics states
extfTdt
rd
2
2
(4.1)
where is the density, r denotes the coordinates of a material point, T is the stress
tensor, and fext is an external volume force. In the stationary case, the equation representing the
force balance is
extfT 0 (4.2)
If the stress tensor is continuous across a stationary boundary between two materials, the
equation can be expressed as:
0)( 121 TTn (4.3)
where T1 and T2 represent the stress tensors of material 1 and 2, and n1 is the outward
normal direction of material 1. The equation shows when material 1 (solid) is surrounded by
material 2 (air/vacuum), the surface force is on the boundary between the materials with the
solid. The user guide of COMSOL [89] deducted the electromagnetic stress tensor expression in
air:
92
TT HBEDIBHDEpIT )2
1
2
1(2 (4.4)
where p is the air pressure, I is the identity 3*3 tensor, E is the electric field intensity, D
is the electric flux density, H is the magnetic field intensity, B is the magnetic flux density. For
the quasi-static magnetic field, the stress tensor can be simplified as:
TBHnBHnTn )()(2
11121 (4.5)
4.5.2 Electromagnetic Model Analysis
In order to study the magnetic force magnitude and distribution in a specific area
according to the equation (4-5), the magnitude and distribution of the magnetic flux density B in
the study area must first be calculated. To perform the simulation, beside the assumptions we
made for the model discussed in Chapter 3, we added two more assumptions:
The material of the steel bar is soft iron instead of Ph-174
The surrounding area is air instead of medium
The meshed geometry of the electromagnet model has 3096 elements and the number of
degrees of freedom is 6251 (Fig. 4.11). The mesh generator partitions the subdomains into
triangular elements, which are allocated based on the geometrical dimensions.
93
Figure 4.11 Meshed Model of an Electromagnet and a Steel Bar
Figure 4.12 Magnetic Flux Density, x component
94
Fig. 4.12 shows the x component distribution with the maximum flux density in a period
at phase zero. In the graph, the magnetic south pole (orange) of the horseshoe electromagnet is
at the lower foot and the magnetic north pole (blue) is at the upper foot. The distribution shows
that the magnetic flux density Bx has greater magnitude in the soft iron core than in the air
space. The maximum B value in the soft iron is 0.221T located at the inner corners of the core.
However the magnitude of Bx dramatically reduces to 0.086T at the boundary of the iron core
and air space. Compared Fig. 4.12 and Fig. 3.2 (a), we can find that the steel bar was
magnetized in the field. The Bx in the area between the magnet and the steel bare increased.
4.5.3 Results
Flux Density Bx VS Distance
X Direction Fig. 4.7 (a) and (b) shows the cross-section plot of magnetic flux density Bn
normal along the X direction (from x=0.08 to x=0.085) at y=0.144 (south pole), and at y=0.159
(north pole), respectively. Compared to Fig. 3.4, the magnitude of the normal magnetic flux
density in Fig. 4.13 dramatically increased when the distance increased to 4mm. The
phenomenon is caused by the magnetizing of the steel bar.
95
(a)
(b)
Figure 4.13 Magnetic Flux Density Bn along x Direction
(a) at Y=0.143; (b) at Y=0.159
96
Figure 4.14 Magnetic Flux Density Bx along y Direction
Bx distributions @ x=0.085
0
0.02
0.04
0.06
0.08
0.1
0.12
0.15
90.
1580.
157
0.15
60.
1550.
1540.
1530.
1520.
151
0.15
00.
1490.
1480.
1470.
146
0.14
50.
1450.
144
y [m]
Mag
net
ic F
lux
Den
sity
, x
com
po
nen
t [T
]
without rod
with rod
Figure 4.15 Comparison of Magnetic Flux Density Bx
97
Y Direction Fig. 4.14 shows the cross-section plot of the x component of the magnetic flux
density along y direction (from y=0.143 to y=0.159) at x=0.085. Fig. 4.14 demonstrates the
same trend as Fig. 3.3, but the magnitude is much greater. Fig. 4.15 compares the difference
between these two plots in absolute terms, which shows the flux density at the edge of the bar is
increased six times. This plot also implies that the force is mostly applied on the edge of the bar,
such that the center part will not be subject to magnetic force. Both Fig. 4.13 and Fig 4.14
illustrate the distribution of the flux density along x and y direction, which directly affects the
calculation of the magnetic force.
Magnetic Force Fx vs. Exciting Current I
The magnetic force acting on the steel bar is calculated by the stress tensor method
described in section 4.6.1. The simulation results show that the magnetic force that the bar is
subjected to along the x direction is 5.1N when the electromagnet is electrified by a 0.4A time-
harmonic sinusoidal current. This is the maximum force the device can provide at 5mm distance
far because the maximum current designed for the device is 0.4A. The force varies as the
exciting current (sinusoidal signal) changes during a period. The relationship between the
magnetic force along x direction and the exciting current is illustrated in Fig. 4.16. In the
previous bio-experiments, the Gelfoam worked as the scaffold, and the stretch force was
normally set as 50%, which means the exciting current was 0.2A. From Fig. 4.16, we can see
the stretch force generated would be about 1.25N.
98
Figure 4.16 Exciting Current vs. Magnetic Force
Magnetic Force Fx VS Distance
Fig. 4.17 shows that the distribution of the flux density is non-linear along the x direction,
which means the magnetic force will also vary along the x direction. The plot demonstrates that
when the bar is closer to the electromagnet, it will be subject to a much stronger force. For a
35mm culture dish, the maximum deformation of the substrate is set to be 2.5mm, so the force
ranges form 1.25N to 4N.
99
Rod_force_x versus Distance
0
10
20
30
40
50
60
0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
Distance[mm]
Ro
d_f
orc
e_x
[N]
Figure 4.17 Distance vs. Magnetic Force
In this thesis, the area around the electromagnet was modeled as air instead of medium, and the
steel bar was modeled by soft iron; more realistic modeling wound give more accurate results.
4.6 Conclusions
In order to validate the performance of the designed apparatus, the effects of uniaxial
cyclic stretch on cell biology have been studied from the aspects of cell proliferation and cell
morphology. The DNA assay results show that the apparatus improved cell proliferation in a
similar way as existing uniaxial stretch devices. The HE staining demonstrated the gross tissue
structure, and the results showed that stretching guided the cell reorganization along the stretch
direction. Moreover, the mechanical parameters (strain distribution, strain rate, and stretch
force) provided by the apparatus have also been investigated. A nondestructive method was
employed to investigate the strain distribution on the surface of the porous scaffold. The results
demonstrate that the strain distribution along the stretch direction is non-uniform, with a high
100
standard deviation. The strain rate was investigated by Photron’s high-speed camera ultima
APX. The results demonstrated the viscoelastic property (time dependent strain) of the scaffold,
and also indicated that the strain rate varies with the strength of the stretch force. The stretch
force was calculated to be 1.25N at 5mm distance from the edge of the electromagnet by
COMSOL Multiphysics, the force was shown to be related to distance and current.
101
Chapter 5
Summary and Future Research
5.1 Summary
Mechanical stimulation is an important regulator in tissue culture for tissue engineering.
Among different forms of mechanical stimuli, uniaxial cyclic stretch is the most commonly
used stimulus for muscle tissues due to its precisely controlled end to end average strain. The
most typical method of applying stretch is motor driven. However, the motor driven method
involves a physical connection between the culture construct and the motor, which requires a
diversity of laboratory devices and increases the risk of contamination.
This thesis proposed a unique idea to design a uniaxial cyclic stretch apparatus based on
non-contact magnetic force. The apparatus developed in this thesis allows researchers to utilize
standard culture dishes, and applies a well-defined strain magnitude to an engineered tissue
patch during culture process, which greatly simplifies the assembly procedure and reduces the
risk of contamination. A prototype of the apparatus has been constructed and tested in a
research lab at the University Health Network in Toronto.
Since the applied mechanical stimulus is generated from magnetic force, the engineered
tissue construct is not only affected by mechanical force, but also exposed to a magnetic field.
102
Thus, the effects of the time-varying magnetic field during the culture process were
investigated. One side effect of using electromagnets, that of a temperature increase, is also
investigated. However, the biomedical experiment results show that neither a weak low
frequency magnetic field (0.1T, 1Hz) nor an increase of 1℃ in temperature has a significant
effect on the cell culture.
In order to verify the performance of the designed apparatus, the effects of uniaxial
cyclic stretch on cell biology have been studied from the aspects of cell proliferation and cell
orientation. The results show that the apparatus improved cell proliferation and cell
reorganization similarly to other known uniaxial stretch devices. Moreover, the mechanical
parameters (strain distribution, strain rate, and stretch force) provided by the apparatus have
also been quantitatively investigated. The strain distribution on the surface of the porous
scaffold along the stretch direction was shown to be non-uniform, with a high standard
deviation. The results of strain rate test indicated that the strain rate varies during stretching due
to the variation of the stretch force. It also proved the viscoelastic property (time dependent
strain) of the scaffold, and provided a new way to monitor the mechanical property of the
engineered tissue during tissue formation. The stretch force, which was calculated using
COMSOL Multiphysics, was shown to be related to distance and current.
5.2 Future Research
This section discusses the future research work that will be pursued after the completion
of the thesis, from the aspects of modeling, bio-experiments, and improvements.
103
5.2.1 Optimizing the Electromagnet Model
The model of the electromagnet discussed in this thesis only considered the quasi-static
electromagnetic field itself, while ignoring the influence of the energy dissipation problems
caused by resistance heating and hysteresis. In the future, the study areas need to cover not only
electromagnetics, but also heat transfer coupled with thermal-electric analysis. The multi-
physical model can help us optimize the relationships among the current, force, and
temperature. Additionally, some of the assumptions need to be revised in order to obtain more
accurate results, such as more specifically characterizing the features of the coil. Moreover, the
model should be upgraded from two-dimensional to three-dimensional.
5.2.2 Implementing Bio-experiments
Although the bio-experiments validated the performance of the apparatus, and
demonstrated that uniaxial stretch can influence cell culture from the aspects of cell
proliferation and cell orientation, the relationship between cell characteristics and the
mechanical parameters are still unclear. The results were affected by multiple parameters such
as stretch frequency, strength, strain, and even stretch pattern and shear flow. The independent
contribution of each parameter is not characterized. Therefore, future experiments should isolate
these mechanical parameters, and systematically examine their independent influence. Such an
experiment would provide researchers an effective platform to explore the mechatransduction of
cells. The upcoming experiment is studying tissue formation by leading the human embryonic
stem cells differentiated to cardiac tissue cells with the defined mechanical stimulation.
104
5.2.3 Calibrating the stretch force
At current stage, the primary controllable parameter of mechanical stimulation in tissue
engineering is strain magnitude, and cultured tissue patches are subjected to predefined
mechanical strain. However, the mechanical force that the patches are subjected to is not fully
studied. According to the simulation results of the COMSOL model, the stretch force is 1.25N
at 5mm away from the magnet, when the magnet is excited by a 0.2A pulsed current. Although
the simulated results of magnetic flux density B are verified by physical experimentation, and
the magnetic force is calculated based on the verified magnetic field, the magnetic force itself is
unverified. Further experiment needs to conduct to determine the stretch force. With the
information of the tested stretch force and filmed deformation, the mechanical properties
(stiffness or elasticity) of the scaffold can be monitored during tissue culture.
5.2.4 Integrating the mechanical and electrical stimuli together
Current bioreactor systems only apply uniaxial cyclic stretch while engineering a tissue
construct. However, the next step of the research will be to combine cyclic stretch and electrical
field stimulation together. The proposed bioreactor is designed to deliver a pulsed electrical
signal to the cell-seeded scaffold with the same frequency as the cyclic stretch. For cardiac
tissue engineering, these external stimuli will mimic the conductive and contractile properties
existing in native heart tissue. Hopefully, these stimuli can lead to better ultrastructural
organization of the engineered tissues.
105
The proposed research aims to develop a novel, compact electrical and mechanical
stimulation apparatus for 3D tissue culture. As shown in Fig. 5.1, the apparatus will consist of a
computerized controller, an electrified mounting tray, and a pair of magnetic metal bars. The
two ends of the cell-seeded scaffold (1) will be fixed by a pair of silver-coated stainless steel
bars (2). The integrated construct (3) will be secured onto the electrified mounting tray (4). The
tray will contain a battery (5) and relevant circuits (6), which will be coated by a biocompatible
material; it will also contain two pairs of metal stoppers (7) serving as electrodes. The mounting
tray will be designed to fit into a standard culture dish (8). Finally, the assembled culture dish
will be mounted between a pair of electromagnets (9). When the electromagnets are electrified
by the controller, the construct integrated with metal bars will experience uniaxial cyclic stretch
between the stoppers and the wall of the dish. As the metal bars touch the metal stoppers, the
construct will also undergo electrical stimulation.
Figure 5.1 Schematic Diagram of the Upgraded Device
106
The device will provide simultaneously and synchronously both uniaxial mechanical cyclic
stretch and electrical stimulus to a 3D engineered-tissue construct, which will subtly mimic the
spontaneous excitation-contraction coupling of native cardiac tissue. With the non-contact
magnetic force and encapsulated battery, the tissue patch will still be cultured in a standard
culture dish. Such an isolated environment will greatly limit the risk of contamination. Using
the new device, researchers will be able to study the effects of both mechanical and electrical
stimuli on cell proliferation, the increased expression of cardiogenesis, and the organization of
the extracellular matrix. Biological experiments will be conducted in the Division of
Cardiovascular Surgery at the University of Toronto. The new apparatus will provide an
effective tool to investigate the mechanism of spontaneous beating of cardiac tissues.
107
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