Effect of mechanical stimulus on cardiac, bone and cartilage tissue

1. EPFL, Tissue Engineering
Semester Paper
Mechanical Stimulation in Tissue Engineering
Authors: Anna Cyganowski, Nadia Vertti, Saurabh Khemka
Professor : Dr. Peter Frey
Spring Semester 2012

2. Contents
1 Introduction 1
1.1 The promise of tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1
1.2 Why mechanical stimulation? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1
2 Current Developments 2
2.1 Cardiac Tissue . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2
2.1.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2
Cardiac Tissue Engineering Principles . . . . . . . . . . . . . . . . . . . . . . . 2
2.1.2 Present State of the Art . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3
Effects of mechanical stretching . . . . . . . . . . . . . . . . . . . . . . . . . . . 3
Effects of mechanical pulsing . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3
Effects of diminished mechanical pulsing . . . . . . . . . . . . . . . . . . . . . . 3
2.2 Bone Tissue . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4
2.2.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4
Bone Tissue Engineering Principles . . . . . . . . . . . . . . . . . . . . . . . . . 5
2.2.2 Present State of the Art . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6
Mechanical stimulation using perfusion based bioreactor . . . . . . . . . . . . . 6
Mechanical stimulation using direct mechanical strain . . . . . . . . . . . . . . 7
2.3 Cartilage Tissue . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8
2.3.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8
Cartilage Tissue Engineering Principles . . . . . . . . . . . . . . . . . . . . . . 9
2.3.2 Present State of the Art . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9
System for studying chondrocytes . . . . . . . . . . . . . . . . . . . . . . . . . 9
Mechanical input . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9
Measurement of cell response . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11
2.3.3 Example of cell engineering mechanotransduction study . . . . . . . . . . . . . 11
3 Future Developments: Microfluidics and Microtechnology 13
4 Conclusions 14
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3. 1 Introduction
1.1 The promise of tissue engineering
Tissue engineering is the final frontier in the field of regenerative medicine. Ever since a group of
MIT researchers presented the infamous Vacanti mouse in 1995, scientists, engineers, clinicians, and
the general public alike have eagerly anticipated the technological advancements that promised to heal
injuries and cure diseases. Instead of relying on an insufficient supply of organ donations or synthetic
materials that integrated poorly with the body’s immune system, patients found hope in the idea that
their own biological cells could be grown into the tissues or organs they needed to survive.
Yet nearly thirty years after the term “tissue engineering” was first used, the technologies that
seemed promising in the lab have yet to make it to the clinic for general use. And, the few technologies
that have been tried, such as synthetic skin for severe burn patients, have severe limits, such as
the lack of sweat glands and hair follicles for cooling the body. Because tissue engineering spans a
significant number of unknowns between biology, medicine, and engineering, this complex field needs
contributions from diverse fields in order to succeed.
1.2 Why mechanical stimulation?
Traditionally, biologists cultured cells by seeding them into an environment that enabled them sur-
vive, with adequate nutrients and temperature. However, in order for cells to develop into viable tissue
for transplantation, they must do more than simply sustain life. Tissue engineers have discovered that
replicating the in vivo environment to which a particular tissue or organ would naturally be exposed
improves the differentiation, organization, and three-dimensional development of cells. Mechanical
stimulation is one factor that is present during cell development, whether during the early fetal stages
or later during the growth, maintenance, and repair of an organ. This is because of a phenomenon
known as mechanotransduction, a process where cells convert mechanical stimuli into electrical or
chemical signals through the cell membrane ion channels. These signals regulate cell processes, such
as the synthesis of growth hormones, which affect cell development into tissue. For these reasons, the
study of mechanical stimulation is key to the advancement of tissue engineered therapies.
Though there is a vast amount of research on the role of mechanical stimulation in tissue engineering,
this report will focus specifically on three tissue types: cardiac, bone, and cartilage.
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4. 2 Current Developments
2.1 Cardiac Tissue
2.1.1 Introduction
In 2008, 1 in 4 deaths in the United States were caused by heart disease, making it the leading cause
of mortality in the country. In addition, over 7 million Americans have suffered from heart attacks, a
staggering statistic that is predicted to worsen as the general population ages [21]. After myocardial
ischemia, the narrowing of the artery causes oxygen depletion and massive localized cell death, which
results in fibrotic scarring and impaired cardiac function. At this point, the heart is unable to repair
itself for two reasons. First, cardiomyocytes, the cells that compose the heart, are limited in their
regeneration potential [3]. Second, the number of myoblasts, the muscle cell type that can divide
and differentiate to form new cells, in cardiac tissue is low [2]. Currently, the most successful clinical
treatment of a failing heart is organ transplantation, but because of a shortage in organ donations as
well as the potential risks of immune reactions, new treatments must be developed.
Cardiac Tissue Engineering Principles
There are several potential methods of myocardial repair currently being studied. One is the direct
transplantation of healthy cells into the damaged areas. To replace necrotic tissue and reduce scarring,
several cell types have been tested, including fetal cardiomyocytes and skeletal myoblasts [12]. The
idea is to graft these cells on top of the wounded areas of the heart to encourage repair. Alternatively,
stem cells from the bone marrow or peripheral blood have been injected into the wounded site to
induce cardiac muscle and vascular tissue growth [1]. However, none of these methods have succeeded
in producing enough myocardial fibers to prevent cell death after engraftment. Instead, the engrafted
myoblasts fail to differentiate and integrate with the cardiac host, which prevents electromechanical
coupling [12]. This lack of synthesis increases the patients risk for more heart problems in the future,
such as arrhythmias.
Tissue engineering has the potential to provide better solutions to this problem. By engineering
cardiac tissue from scaffolds or biomaterial gels seeded with cells, an engineer can better control
the parameters of the engineered constructs, such as shape, size, development, and function [12].
Ultimately, one goal of cardiac tissue engineering is to develop a cell sheet patch that can be grafted
to the heart to stimulate regeneration. However, the main challenge is creating a muscle that is
robust while reducing the risk for necrosis after transplantation. Zimmerman et. al. proposed that
a successful cardiac tissue engineered construct should meet the following five criteria: contractility,
electrophysiological stability, flexibility and mechanical robustness, vascularization or the ability to be
quickly vascularized in vivo, and non-immunogenicity [23]. To date, engineered tissues have not yet
met all of these criteria.
Among the techniques used to improve the differentiation of cardiac tissue, many have the same
goal in mind: to mimic the complex environment of the developing heart in vivo. Though there
are many factors that contribute to this, such as biochemical signaling, the mechanical stimulation
experienced by cardiac tissue is perhaps the most obvious. Pumping over 40 million times a year
at 60-100 beats per minute [23], the heart experiences a significant amount of stresses and strains.
Studying the boundary conditions of this mechanical stimulation as experienced by fetal tissue would
not only expand scientific knowledge in the field of developmental biology, but would also give tissue
engineers the tools they need to replicate this phenomenon in vitro.
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5. 2.1.2 Present State of the Art
Effects of mechanical stretching
Many researchers have incorporated the use of mechanical stimulation for cardiac tissue engineering,
primarily in bioreactor environments. Some of the earliest and most cited work has been done by
Zimmerman et. al. at the University of Hamburg in Germany. This group created biomaterial
constructs of a combination of collagen type I, extracellular matrix proteins, and freshly isolated heart
cells from neonatal rats and embryonic chickens. The advantage of using this mixture is that it can
be casted into molds of various shapes, which allows for more experimental structures, ranging from
lattice to circular geometries [4]. After culturing for 5-7 days, the developing tissue solidified and
began to contract. To further enhance development, the construct continued to culture with added
cyclic mechanical stretch for 5-10 more days.
The group found significant evidence to suggest that mechanical stimulation improved the develop-
ment of the engineered heart tissue. First, the morphology of the stretched tissue was more advanced,
with denser regions of more developed and longitudinally oriented rod-shaped cells. The cross-sectional
diameter of the stretched cells were 41% greater than those without stretching [4]. The contractile
forces of the cells after mechanical stimulation were between 2-4 times higher than the control cells,
both at rest and during stimulation with calcium. Fig. 2.1 shows the experimental setup, including
the culture dish with engineered heart tissue structures before stretching, the motorized stretching
device, and the schematic representation of the phasic stretch used at 1.5 Hz. While this work was
useful in showing that mechanical stimulation is beneficial to the development of engineered tissue,
the results have their limits. The engineered tissue constructs in this case are only 0.4 millimeters,
which is too thin to use in grafting applications.
Effects of mechanical pulsing
Another method of introducing mechanical stimulation to tissue culture is through the use of biore-
actors. Simon Hoerstrup et. al developed an in vitro pulse duplicator system to focus on a different
aspect of cardiac tissue engineering. Instead of developing cardiac tissue, the goal of this group was
to engineer heart valves in vitro as a substitute for the mechanical heart valve devices used in the
clinic today. Though heart valve devices give patients a practical alternative to organ transplants – an
option that is unavailable for those who have cardiac tissue damage – these mechanical devices often
fail in 30-35% of patients within 10 years [8] . Some problems caused by mechanical valves include an
increased risk of thromboembolism and tissue deterioration. Especially in young and growing children,
the durability and size limitations of mechanical valves are not ideal. Previous tissue engineered heart
valves are also structurally immature and mechanically weak. To improve the development of heart
valves in vitro, these constructs were cultured in a bioreactor for 14 days, with flow and pressure
conditions gradually increasing. The flow ranged from 125 mL/min to 750 mL/min and the pressure
increased from 30 mm Hg to 55 mm Hg. Afterwards, cardiopulmonary bypass was used to implant
the engineered constructs into lambs and later removed after 1 day and 4, 6, 8, 16, and 20 weeks [8].
The results of this experiment were encouraging. After 20 weeks of implantation in the lambs,
the heart valve tissue constructs were found to be mechanically comparable to native tissue. The
extracellular matrix and DNA levels were also similar to native tissue at that point in development.
However, there is a great need for more future work before this technology can be transferred to the
clinic. Though the engineered heart valve was successful in vivo at 20 weeks, the longer term fate is
unknown. In addition, since the number of implants in this experiment was small, more tests need
to be done to increase statistical significance. Other improvements to this work include creating a
scaffold that is more accurate in shape, finding the optimal cell source for seeding, and optimizing the
pressure loading conditions in the bioreactor to simulate the biological environment.
Effects of diminished mechanical pulsing
A more improved bioreactor design has been implemented by Yuji Narita et. al., which uses a pulse
duplicating system that is capable of softening the usual peaky pressure waveforms seen in mechanical
3

6. Figure 2.1: Setup used by Zimmerman et. al. to stretch engineered heart tissue constructs [4]
stimulation [14]. Fig. 2.2 shows the general structure of this bioreactor, which includes a balloon
chamber with a mechanical valve and a compliance chamber that softens the peaks of mechanical
stimulation. Perviously, similar bioreactors for mechanical stimulation included a compressed air res-
pirator with a peristaltic pump, and couldn’t handle various types and shapes of scaffold materials.
Because of the balloon chamber, the tissue constructs could be exposed to both shear and stretch
stresses, and a wide range of pulsatile flows. This also means that the pressure profile resembled
more closely the kind seen in vivo during natural development. The results of this experiment clearly
demonstrated an improvement: the cell counts, protein content, and pproteo-glycan-glycosamino gly-
can levels of tissue constructs grown in this bioreactor was higher than those in both the static and
traditional ”peaky pressure” stimulation chambers [14].
2.2 Bone Tissue
2.2.1 Introduction
Bone, a tissue with remarkable rigidity, provides structure to the body. Bone not only supports
and protects the internal part of the body, but also produces blood cells and stores minerals and
growth factors. Apart from traumatic accidental injuries, there are several other common bone defects
that need treatment, such as osteomyelitis (bone or bone marrow infection), osteogenesis imperfecta
(genetic bone disorder), and osteoporosis (brittle bone disease).
Osteoarthritis (degenerative joint disease) alone is found to affect 26.9 million US adults as of 2005
with more than 780,000 total joint replacements in 2006. In 1997, total expenditure for knee and
joint replacement was 7.9 billion dollars. Currently, the standard procedure for critical-sized bone
4

7. Figure 2.2: Pulse duplicating bioreactor with the ability to avoid peaky pressure waveforms [14]
defects is autogenous grafting, in which bone from another site is taken and grafted. However, higher
complications, pain at the donor site, increased risk of infections, and restricted availability limits
its application in traumatic bone injuries. Alternatively, the use of bone tissue from other humans,
known as allografting, is being employed, but the risk of infection transfer, host immune response, and
limited availability are associated concerns. Xenografting, the grafting of bone from a non-human, is
considered to be unsuitable due to increased risk of disease, virus transmission, host rejection, and
immunogenicity.
The aforementioned problems associated with current procedures limits the supply of the graft while
the demand increases year by year, forcing scientists, engineers, and clinicians to look for alternatives.
Tissue engineering seems to be a promising in producing new bone tissue by utilizing the body’s
natural response to tissue damage.
Bone Tissue Engineering Principles
Tissue engineering aims to built the organs in vivo, ready to be transplanted to patients in need.
To grow living tissue in vitro, cultured cells are seeded on bioactive degradable scaffolds that pro-
vide the physical and chemical cues to guide their differentiation and assembly into three-dimensional
(3D) tissues. An ideal bone scaffold should present a physiochemical biomimetic environment, pro-
vide temporary mechanical support to the affected area, contain a porous structure allowing bone cell
migration and vascularization, and deliver drugs in a controlled manner to enhance rapid healing. [16]
In a scaffold, cells are seeded and allowed to proliferate. Osteoblasts have been widely used as these
cells produce bone extracellular matrix (ECM) in the scaffold; however, their ability to proliferate is
significantly less than their precursor, mesenchymal stromal cells (MSCs). MSCs are mostly isolated
from bone marrow aspirates and have the ability to differentiate into diverse mesenchymal lineages,
including osteoblasts, chondrocytes, adipocytes, and myocytes. [15] One major problem with this
approach of in vitro culturing of cells in scaffolds to produce bone grafts is the long time required to
create the scaffold, which forces the patient to wait. In order to enhance growth, various techniques
are being employed including the use of growth factors and bioreactors that provide mechanical stim-
ulation. This section of the report will focus on the effect of mechanical transduction on bone tissue
engineering.
The inferior mechanical properties of tissue engineered constructs has driven researchers to parse
the role of mechanotransduction. This process converts mechanical stimuli into biochemical signals,
leading to the adaption of bone to the mechanical loading. One of the most important mechanical
stimuli is the shear stress produced by the blood flow in bone porous space. For humans, the level
of shear stress is in the range of 0.8-3 Pa. Moreover, the muscle contraction and body movements
result in continuous exposure of bone to mechanical stimuli by changing hydrostatic pressure, direct
cell strain, fluid flow-induced shear stress, and electric fields.[16] It has been demonstrated that fluid
induced shear stress enhances the intracellular Ca2+ release in vascular endothelial cells, activating
5

8. Figure 2.3: apparatus used for Diagram of mechanical strain (A) and fluid flow (B) experiments.
[20]
other proteins of signaling pathways. Moreover, it has been suggested that fluid flow up regulates the
expression of many signaling molecules, paracrine factors (example- prostaglandin E2) and mRNA
of bone matrix proteins (e.g., osteopontin). Different growth factors, such as the insulin/like BMPs,
have also been found to play role in mechanical stimulation. In recent years, various bioreactors have
been designed to mimic the in vivo environment by facilitating the proper nutrient supply across the
3D scaffold, cell growth, and mechanical stimulation. In the following section, we explore the present
state-of-art in bone tissue engineering and role of mechanical stimulation.
2.2.2 Present State of the Art
Mechanical stimulation using perfusion based bioreactor
In early 1997, R. Smalt and colleagues compared the ability of direct mechanical strain and fluid flow
induced shear stress to stimulate the bone cell growth. Rat calvarial osteoblast cells were extracted
and suspended in adequate buffer. For the mechanical strain testing, the cells were added to wells on
a tissue culture-treated strip of polystyrene film followed by the attachment to metals bars capable of
moving relative to each other by the electromagnetic force (Fig.2.3). While in flow experiments, glass
slides were mounted on a parallel plate flow chamber, and HEPES-buffer with 0.1 % bovine serum
albumin was allowed to flow in controlled manner. Shear stress of the range 0-80 dyn/cm2 was applied
(Fig.2.3). They observed that although the bone cells were unresponsive to mechanical strain up to
5,000 µ-strain, both primary cultures containing osteoblastic cells and osteoblastic cell lines responded
to even low levels of wall-shear stress, resulting in higher production of PEG2. It was shown that
fluid flow shear stress in the range of 8-30 dyn/cm2 is more likely to be the stimulus acting on bone
cells in vivo. [20]
Although the above mentioned experiment demonstrated the ability of flow induced shear stress
in stimulating cell growth, the experiment was limited by the use of monolayer culturing. This is
quite different from the environment in vivo, which is 3-dimensional. Perfusion bioreactors have been
designed to achieve the flow induced shear stress with low rate of perfusion (e.g., 0.01-3 mL/min)
to provide adequate diffusion of nutrients through 3D scaffolds and necessary shear stress for cell
proliferation. By mimicking the physiological loading of bones, which is cyclic in nature (e.g. the
direction of fluid flow in stance phase of gait is different from the swing phase of the gate), Vance et
al. in 2005 used perfusion bioreactors with oscillating fluid flow. They used the mouse osteoblastic
6

9. cell line MC3T3 cultured in standard MC3T3 medium, which were statically seeded onto the calcium
phosphate scaffolds having the porous structure similar to trabecular bone. The bioreactor in which the
seeded scaffold were loaded comprised a syringe pump providing the constant low perfusion flow rate
with a linear actuator to generates a larger magnitude flow (mechanical stimulus) and an oscillatory
flow profile apart from other parts of bioreactor (Fig. 2.4). They observed that PGE2 levels increased
Figure 2.4: (A) Bioreactor system: a linear actuator and syringe pump are used to drive culture
medium through the bioreactor chambers. Flow rates and profiles are monitored with a
flow probe. (B) Bioreactor flow chamber viewed in longitudinal section. A single calcium
phosphate scaffold seeded with cells. [22]
by 2.5-fold and 4.5 folds compared from static controls, to an average fluid-induced shear stress of
0.0007 Pa from perfusion flow and 30 min of oscillating flow at 1 Hz, producing a shear stress of
approximately 1.2 Pa, interposed on the same perfusion flow respectively. [22] These findings clearly
demonstrate the role of mechanical stimulation in enhancing the proliferation of osteoblastic cells.
Recent advances include use of decellularized bovine cancellous bone (cylindric, 4X4 mm) as a
scaffold seeded with human MSCs to mimic biological topography in an novel perfusion bioreactor
capable of cultivating six tissue constructs simultaneously. By increasing the flow rate from 100 mm/s
to 400 mm/s increment in uniform cell distribution, Collagen growth, osteopontin, bone sialoprotein
(BSP2) was observed. [6]Similarly in 2008, Jagodzinski et al. observed the increase in cell proliferation
and osteocalcin (OC) level increment in above mentioned scaffold of size 4x20mm using a custom made
perfusion based reactor with cyclic compression. [10]. These finding clearly demonstrate the role of
mechanical stimulation in enhancing the proliferation of osteoblastic cells, different growth factors and
signaling molecules.
Mechanical stimulation using direct mechanical strain
In 1892, Julius Wolff depicted the role of mechanical loading in bone remodeling. Although the
exact mechanisms are still unknown, Wolff found that loading the bone with various strains affected
bone loss, growth, and maintenance. Strains below 500 µ-strain resulted in bone loss, while loading
up to 1000 µ-strain maintained the original bone geometry and mass. Meanwhile, strains between
1000 - 4000µ-strain increased new bone formation progressively. [17]
Various type of direct mechanical strain include stretching, bending, contraction or compression,
and few of them will be discussed below. The use of four point bending (Fig. 2.5) in the study
7

10. of human bone marrow stomal cells (hBMSCs) seeded in 3-D partially demineralized bone scaffold
has been shown to increase osteogenic differentiation. Increased levels of alkaline phosphatase (ALP)
activity (around 82–90 %), mineralized matrix production, and gene expression of ALP (by 218-257 %)
and OP was also observed. The bioreactor utilized the tensile strain in the range of 44–2151 µ-strain
and compressive strain of -51 to -2342 µ-strain. In this bioreactor, the force of the linear actuator is
transmitted to four-point bending of the construct via the transduction beam during loading period.
[13] Similarly in 2004, Ignatius et al. reported the effect of cyclic uniaxial mechanical strain (Fig.
2.5) of 10,000 µ-strain on the human osteoblastic precursor cell line(hFOB1.19) in a type I collagen
scaffold. They observed the high proliferation of the osteoblastic cells along with higher gene expression
of H4 (a gene related to proliferation), osteoblastic differentiation and matrix production during the
cultivation period of 3 weeks. Further, the growth of ECMs were found in the direction of mechanical
strain applied. [9] Although the existing bioreactor systems for direct mechanical stimulation have
Figure 2.5: (A) illustrates the four-point bending method. On the left, the scaffold is shown in the
passive state. On the right, mechanical loading is applied. (B) demonstrates the principle
of uniaxial cyclic stretching, which is applied in elastic silicon dishes. Cells are embed-
ded in three-dimensional collagen type I matrices (dark gray). (C) represents a uniaxial
mechanical loading device. A plunger is pushed in a cyclic manner on the scaffold.
shown beneficial effects on proliferation, osteogenic differentiation, and matrix formation, the use of
bio-mechanically unstable Collagen I gels as matrices for mechanical strain-based cellular stimulation
is disadvantageous where just a certain stability is required for the initial implantation and no further
stability is needed. [17]
2.3 Cartilage Tissue
2.3.1 Introduction
Cartilage is a type of connective tissue in the body. It is made of cells called chondrocytes em-
bedded in a matrix, and strengthened with fibers of collagen, an abundant ground substance rich in
proteoglycan and sometimes elastin. It is an important tissue because of its role as a template for the
osseous skeleton, a regulator of skeletal growth, the articular covering for bone ends. There are three
different types: hyaline cartilage, elastic cartilage, and fibrocartilage. Cartilage is avascular, meaning
that it is not supplied by blood vessels; instead, nutrients diffuse through the matrix, helped by the
pumping action generated by compression of the articular cartilage or flexion of the elastic cartilage.
Cartilage is usually flexible, depending on the type. Some of the bodily structures that are composed
8

11. of cartilage include the ears, nose, ribcage, and intervertebral discs.
According to the Department of Bioengineering of Rice University in the United States, the po-
tential benefits of cartilage regeneration and replacement therapies are enormous. Tens of millions of
Americans suffer from acute trauma to musculoskeletal tissues as well as various degenerative cartilage
conditions. Currently, patients have little recourse beyond surgical techniques to ameliorate the symp-
toms of disease or restore some function to damaged tissue, and cannot otherwise address the issue
of healing articular cartilage. It is also important to mention the research around the consequence of
the application of mechanical stimuli to the mesenchymal stem cells to compromise into chondrocytes
that is carried on around the world by diverse institutions. Much interest is put into the cartilage
tissue development.
Cartilage Tissue Engineering Principles
In the tissue engineering domain, cartilage can be viewed as a biological tissue in which cells within a
dense extracellular matrix (ECM) are presented with a complex combination of physical forces, flows,
and biological signaling factors. From the mechanical point of view, the articular cartilage (hyaline
and fibrocartilage type) is subjected to a range of static and dynamical mechanical loads because of
its presence in human synovial joints. These loads can occur at a high frequency (short-duration) and
at long term, even static conditions, within the physiological range. These forces can be applied in a
compressive, tensile and shear way. The values of stress that cartilage receives can even reach 10 to
20 MPa in some physical activities[7].
In order to understand how chondrocytes form a tissue with the needed material properties, as well
as the ECM molecular structure for performing its physiological function in the human body, their
mechanisms must be understood first. This is the feedback process between physical stimuli that the
cells receive and their responses (both by multiple regulatory pathways) at the molecular, cellular and
tissue levels. As a consequence of the mechanical stimuli in their microenvironments, chondrocytes
can synthetize, assemble, and degrade proteoglycans (PGs), collagens, glycoproteins and other matrix
molecules. One important area of study is determining how these mechanical inputs affect developing
cartilage [19].
2.3.2 Present State of the Art
System for studying chondrocytes
Since the study of the mechanism by which chondrocyte respond to mechanical stimuli is difficult
in vivo, there has been a development of in vitro models. These in vitro models have changed through
the years of cell cultivation history. Three dimensional chondrocyte/gel culture systems are impor-
tant nowadays because they can preserve or emulate native tissue structure and enable quantitive
correlations between mechanical and biological parameters. Also as a consequence of delimitation of
space, cell matrix interactions and chondrocyte gene expression can be preserved in these systems.
The comprehension of the relation ECM-Chondrocytes is crucial for the tissue engineering of cartilage.
Some of the hydrogels utilized by researchers for doing so are agarose, alginate and polyglycolic acid
(PGA). In these substrates cells can be cultivated and then scientists look at chondrocyte phenotypic
expression, proliferation, and accumulation of a PG-rich ECM during long periods of time[7].
Mechanical input
Researchers have studied the effects of various types of mechanical stimulation on cartilage cells,
including applied mechanical compression (load or displacement control), hydrostatic pressure, physic-
ochemical stimuli (pH and osmolarity), and electrical currents. This is because the mechanical inputs
that the cartilage receives is a combinations of mechanical stresses simultaneously during joint motion
on an intermittent basis. And among these types of stimuli, there have also been a large number of
experiments with different ranges of stimuli inputs over a broad variety of frequencies to the cells in-
volved. In general, it has been found that the metabolic response to compression in vitro shows similar
9

12. trends to those seen in animal studies. These suggest that static compression would significantly in-
hibit the synthesis of PGs and proteins, whereas dynamic compression can markedly stimulate matrix
production, depending on the amplitude and frequency[19].
How is it possible to apply mechanical stimulation? Experiments have shown that compression of
cartilage causes deformation of cells and matrix, while hydrostatic pressure gradients create intersti-
tial fluid flow. Fluid convection and separation of counterions from the fixed charge groups of the
proteoglycan constituents gives rise to electrical streaming potentials and currents. In contrast, tissue
shear deformation of a poroelastic tissue does not induce volumetric changes, intratissue fluid flow,
or pressure gradients. All these effects of the mechanical stimuli can be modeled thanks to the devel-
opment of computational simulations, which give researchers a framework for what to expect when
they apply the mechanical stimulation to cells. The different types of mechanical stimuli are shown
in figure 2.6.
Figure 2.6: Schematic overview of stimulation of chondrocytes cultures with examples of bioreactors
that can be used. [18]
10

13. Measurement of cell response
Another technique in cartilage tissue engineering is the measurements of cell deformation, which
has been used to indirectly determine the material properties of chondrocytes. These measurements
are mostly performed within compressed gels in three ways: micropipette aspiration, cytoindentation
and atomic force microscopy (Figure 2.7).
Figure 2.7: Illustration of the three direct techniques currently used to characterize the biomechanical
characteristics: (A) micropipette aspiration, (B) cytoindentation and (C) atomic force
microscopy. [19]
The majority of data on the properties of chondrocytes has been obtained using the micropipette
aspiration technique. A long range of measurements of the Young’s modulus and Poissons ratio of
chondrocytes has been performed, as well as volumetric properties and viscosity. These values vary due
to the experimental differences in the cells and scaffolds chosen, the stresses applied, and the method
of measurement. Measuring the rate at which chondrocytes can sense and respond to mechanical
stimuli can give valuable insight into intracellular regulatory mechanisms.
Finally, the effects of various mechanical forces on gene expression are investigated. For example,
static and intermittent hydrostatic pressure increased the expression of transforming growth factor
β, as well as aggrecan and type II collagen mRNA, in high-density monolayer cultures. In isolated
human chondrocytes grown in monolayer, constant fluid shear forces stimulated expression of mRNA
for tissue inhibitor of metalloproteinase. Dynamic mechanical forces have also been shown to influence
matrix gene expression.
2.3.3 Example of cell engineering mechanical stimulation study
Tissue engineering of cartilage can have unusual applications such as the study of this tissue grow in
space [5]. This interest is mainly applicable for long term space flights, for passengers who are exposed
to microgravity as well as launch and landing events. Results from these studies have enriched the
understanding of the effects of pseudo-weightlessness on prolonged immobilization, hydrotherapy, and
intrauterine development. These studies were conducted on the Mir Space Station and the results were
compared to those obtained by the same system on Earth. Specifically, three-dimensional cell-polymer
constructs consisting of bovine articular chondrocytes and polyglycolic acid scaffolds were grown in
rotating bioreactors, first for 3 months on Earth and then for an additional 4 months on either
Mir (10−4-10−6 g) or Earth (1 g) as shown in figure 2.8. Chondrocytes were seeded onto scaffold
constructs (5 million cells per PGA disc) in spinner flasks stirred at 80 rpm in a 37 ◦C humidified
10% CO2 incubator. After 6 days, cell–PGA constructs were transferred into rotating bioreactors
configured as the annular space between a 5.75-cm diameter polycarbonate outer cylinder and a 2-
cm diameter hollow inner cylinder. The entire vessel was rotated as a solid body around its central
axis while gas exchange was provided by pumping incubator air through the inner cylinder. Culture
11

14. medium was replaced at a rate of 50% every 3-4 days, and the vessel rotation speed was gradually
increased from 15 to 28 rpm over 3 months, to induce mixing by gravitational construct settling. After
3 months, constructs were transferred into each of two flight-qualified rotating, perfused bioreactors
(BTS) for an additional 4 months of cultivation on either the Mir Space Station or on Earth. Both
environments yielded cartilaginous constructs, each weighing between 0.3 and 0.4 g and consisting of
viable, differentiated cells that synthesized proteoglycan and type II collagen. Compared with the
Earth group, Mir grown constructs were smaller, more spherical, and mechanically inferior, as shown
in figure 2.9. [5]
Figure 2.8: Experimental design of the bioreactor.[5]
Figure 2.9: Transmission electron micrograph of constructs from (A) Mir and (B) Earth. [5]
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15. 3 Future Developments: Microfluidics and
Microtechnology
The emergence of microfluidics has inspired new strategies to more easily induce cellular mechan-
otransduction [11]. Cellular stimulation by locally varying fluid shear can serve to accurately alter
membrane surface tension as well as produce direct compressive and strain forces onto cells. More-
over, microtechnology has increased the possibilities for individual cell-level actuation and readout.
Microsystems are perfectly suited for the analyses of individual cells, as well as more complex cell
culture and tissue systems, since the integrated tools match the length scales of cells (550 mm).
Since small liquid volumes be controlled precisely, the cellular environment can be controlled quite
accurately. Microsystems technology also offers the integration of microsized, movable, fast-operating
components, such as valves , pumps, and microsized cell culture chambers. Figure 3.1 shows these
sort of features:
Figure 3.1: Applications of µfluidic devices for mechanotransduction [11]
Shear forces: The laminar flow regime that is achievable with microdevices facilitates the application
of well-defined flow rates and profiles within the microfluidic channels and hence, allows for the local
stimulation of cells with precise shear forces.
Tensile stress: A thin membrane separating air-filled epithelial from the liquid-filled endothelial
compartments could be stretch-activated by adjacent vacuum channels, leading to a cyclical mechanical
strain on the indicated cell types.
Cell compression: Microfluidic cytometers have been developed to identify potentially cancerous
cells based on their compression responses in a high throughput manner. By combining a polymer
cantilever-based method with micro restriction flow through, the cells mechanical response to physical
compression could directly be measured.
Intracellular architecture: Micropost arrays have also been successfully employed in more sophisti-
cated designs. Clamping a stretchable polymeric membrane possessing a micropillar fabricated surface
into a cell-culturing module facilitated the measurement of intra-cellular microscopic forces with dy-
namic extracellular microscopy.
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16. 4 Conclusions
Mechanical stimulation plays an important role in the culture of cells, especially those that are
naturally under mechanical stress in the human body. When applied to cultivating cells, these inputs
start a cascade of signals that influence the development of the tissue. The relationship between the
extracellular matrix and cell development, particularly the mechanical feedback process between them,
is important for tissue engineers to understand in order to induce adequate mechanical stimulation
to out-of-body environments. The many experiments that have involved mechanical stimulation have
become a framework for upcoming studies. This report describes some but not all of them, since many
factors are involved.
The long term goal of tissue engineering is to develop in vitro tissues and organs with properties and
behaviors closely resembling those developed naturally in living beings. The behavior and development
of cells provide vital clues to reaching this goal. Developing tools that allow researchers to better
manipulate cells and mimic in vivo environments is a crucial step in the growth of this field. Though
current bioreactors successfully apply mechanotransduction to induce cell proliferation and grow 3D
grafts, more research needs to be done to optimize the transplantation process. In addition, though
many present-day bioreactors provide adequate physical, chemical, and biological environments, a
further optimization of specific stress, load, and biophysical stimulus combinations is needed to refine
growth processes according to the needs of specific cell types.
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