coupling of mechanical and electromagnetic fields
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Coupling of Mechanical and Electromagnetic Fields Stimulation Coupling of Mechanical and Electromagnetic Fields Stimulation
for Bone Tissue Engineering for Bone Tissue Engineering
Alyaa I. Aldebs Wright State University
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COUPLING OF MECHANICAL AND ELECTROMAGNETIC FIELDS
STIMULATION FOR BONE TISSUE ENGINEERING
A thesis submitted in partial fulfillment of the
requirements for the degree of Master of Science
in Biomedical Engineering
By
ALYAA I. ALDEBS
B.Sc. Biomedical Engineering, University of Al-Nahrain, 2012
2018
Wright State University
WRIGHT STATE UNIVERSITY
GRADUATE SCHOOL
April 27, 2018
I HEREBY RECOMMEND THAT THE THESIS PREPARED UNDER MY
SUPERVISION BY Alyaa I. Aldebs Entitled Coupling of Mechanical and Electromagnetic
Fields Stimulation for Bone Tissue Engineering BE ACCEPTED IN PARTIAL
FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF Master of Science
in Biomedical Engineering.
________________________________
Jaime E. Ramirez-Vick, Ph.D.
Thesis Director
________________________________
Jaime E. Ramirez-Vick, Ph.D.
Department Chair
Committee on Final Examination
______________________________
Jaime E. Ramirez-Vick, Ph.D.
______________________________
Ulas Sunar, Ph.D.
______________________________
Nasim Nosoudi, Ph.D.
_______________________________
Barry Milligan, Ph.D.
Interim Dean of the Graduate School
iii
ABSTRACT
ALDEBS, ALYAA I. M.S.B.M.E. Department of Biomedical, Industrial and
Human Factors Engineering, Wright State University, 2018. Coupling of
Mechanical and Electromagnetic Fields Stimulation for Bone Tissue Engineering.
Alternative bone regeneration strategies that do not rely on harvested tissue or exogenous
growth factors and cells are badly needed. However, creating living tissue constructs that
are structurally, functionally and mechanically comparable to the natural bone has been a
challenge so far. A major hurdle has been recreating the bone tissue microenvironment
using the appropriate combination of cells, scaffold and stimulation to direct
differentiation. This project presents a bone regeneration formulation that involves the use
of human adipose-derived mesenchymal stems cells (hASCs) and a 3D scaffold based on
a self-assembled peptide hydrogel doped with superparamagnetic nanoparticles (NPs).
Osteogenic differentiation of hASCs is achieved through the direct stimulation by
extremely-low frequency pulsed electromagnetic fields (pEMFs) and the indirect
mechanical stimulation, through NP vibration induced by the field. This 3D construct was
cultured for up to 21 days and its osteogenic capacity was assessed. Cellular morphology,
proliferation, viability, as well alkaline phosphatase activity, calcium deposition were
monitored during this time.
The results show that the pEMFs have no negative effect on cell viability and
induce early differentiation of hASCs to an osteoblastic phenotype when compared to a
iv
cell without biophysical stimulation. This effect results from the synergy between
the pEMF and NP that acts as remote stimulation of the mechanotransduction pathways
which activate biochemical signals between cells to go under differentiation or
proliferation. The use of this approach offers a safe and effective treatment option for the
treatment of non-union bone fractures. In addition, this formulation can be directly injected
into the wound site, making it minimally invasive as well.
v
Table of Contents
ABSTRACT ...................................................................................................................... iii
Table of Contents .............................................................................................................. v
List of Figures .................................................................................................................. vii
Acknowledgment .............................................................................................................. ix
CHAPTER I ...................................................................................................................... 1
Introduction ....................................................................................................................... 1
1.1 Project overview ........................................................................................................ 1
1.2 Objectives .................................................................................................................. 5
CHAPTER II ..................................................................................................................... 7
Literature Review ............................................................................................................. 7
2.1 Bone structure ........................................................................................................... 7
2.2 Bone anatomy ............................................................................................................ 7
2.3 Bone composition .................................................................................................... 11
2.3.1 Bone Extracellular Matrix ................................................................................ 11
2.3.2 bone cells .......................................................................................................... 14
2.3.3 growth factors ................................................................................................... 17
2.4 Bone diseases .......................................................................................................... 19
2.5 Bone replacement and tissue engineering ............................................................... 20
2.6 Cellular sources for bone tissue engineering........................................................... 22
2.7 Three-dimensional scaffold for bone tissue engineering..……..………………25
2.7.1 Self-assembled peptide scaffold ....................................................................... 26
2.8 Biophysical stimulation ........................................................................................... 30
2.8.1 Effect of mechanical stimulation ...................................................................... 30
2.8.2 Effect of electromagnetic fields ........................................................................ 34
vi
2.9 Effect of magnetic scaffold under EMF stimulation for bone regeneration ........... 38
CHAPTER III ................................................................................................................. 46
Materials and Methods ................................................................................................... 46
3.1 Cell culture .............................................................................................................. 46
3.2 Three-dimensional cells encapsulation and gel formation .............................. 46
3.3 Cells viability assay ................................................................................................. 47
3.4 Cells differentiation assay ....................................................................................... 48
3.5 Measurement of mineralization ............................................................................... 49
3.6 Cells morphology .................................................................................................... 49
3.7 Alkaline phosphatase staining ................................................................................. 50
3.8 Fourier transform infrared spectroscopy (FTIR) analysis ....................................... 51
3.9 Electromagnetic fields exposure system ................................................................. 51
3.10 Statistical analysis ................................................................................................. 53
CHAPTER IV.................................................................................................................. 54
Results &Discussion ........................................................................................................ 54
4.1 Results & Discussion ..................................................................................... 54
4.1.1 Cell viability ............................................................................................ 54
4.1.2 Differentiation to osteoblasts ................................................................... 59
4.1.3 Mineralization .......................................................................................... 63
4.1.4 Cell morphology ...................................................................................... 65
4.1.5 Alkaline phosphatase staining .................................................................. 68
4.1.6 FTIR analysis ........................................................................................... 70
4.2 Conclusion ............................................................................................................... 71
References .............................................................................................................. 74
Appendix I ....................................................................................................................... 89
vii
List of Figures
Figure Page
1. Three-dimensional structure of bone that shows the cortical and trabecular
components…………………………………………………………………........10
2. Topographic relationships of different bone cells that initiate from different
origins…………………………………………………………………….……...15
3. Amino acids sequence of self- assembled peptide (RADA-16) ………………...28
4. The mechanosensors that resulted from changes in ions channels, proteins, and
cytoskeleton to activate the intercellular signals…………………………….......32
5. Schematic illustration of using Helmholtz coil to generate pEMF……………...35
6. Electromagnetic exposure system: a) closed mu-metal box and b) Helmholtz coils
with cells.…………………………………………………………………..……53
7. LDH assay. The proliferation of hASCs seeded within self-assembled peptide
hydrogel with and without NPs under extremely low-frequency pEMF of 1 mT
was quantified, after 7, 14, 21 d. *p < 0.05; **p < 0.001; ***p < 0.0001, indicate
statistically significant differences between basal media group (control) and the
other groups. The mean values are calculated from the average results of three
samples, the results are represented as mean ± SD..…………………….……….56
8. Alkaline phosphatase activity. The differentiation of hASCs cells was assessed at
7, 14, 21 d. ALP values were normalized with the number of cells of each sample.
*p < 0.05; **p < 0.001; ***p < 0.0001, indicate statistically significant
differences between basal media group (control) and the other groups (one-way
ANOVA followed by Dunnett test). The mean values are calculated from the
average results of three samples, the results are represented as mean ±
SD………………………………………………………………….………….…62
9. Mineralization assay. Calcium concentrations were quantified after 7,14, 21 days.
Error bar represents the s.d. p < 0.05, indicate statistically significant differences
between MSCs media gel+cells group (control) and the other groups (one-way
ANOVA followed by Dunnett test). The mean values are calculated from the
average results of three samples, the results are represented as mean ±
SD…………………………………………………………………………..……65
viii
10. Phalloidin-labeled actin filaments stain (red) and DAPI stain (blue) for hASCs
within hydrogel at 3 and 14 days. A) basal media, B) basal media + NPs, C)
osteogenic media, D) osteogenic media + NPs, E) basal media + pEMF, F) basal
media + NPs + pEMF, G) osteogenic media + pEMF, H) osteogenic media + NPs
+ pEMF. Scale bars at day 3 and 14 are 100 µm (10X) and 50 µm (20X),
respectively. Biochemical stimulation – osteogenic induction
media…………………………………………..……………………………...… 67
11. Alkaline phosphatase stain for hASCs cells seeded within hydrogel at day 14,
scale bar = 200 µm (10X). A) basal media, B) basal media + NPs, C) osteogenic
media, D) osteogenic media + NPs, E) basal media + pEMF, F) basal media +
NPs + pEMF, G) osteogenic media + pEMF, H) osteogenic media + NPs +
pEMF. Biochemical stimulation – osteogenic induction media . ……..………..69
12. FTIR analysis of self-assembled peptides with and without iron oxide
nanoparticles……………………………………………………………………..71
ix
Acknowledgment
In the beginning, I would like to express my sincere thanks to my thesis advisor, Dr.
Ramirez-Vick for his support during my research time and for advising me to find the
answer to all my concerns through his valuable comments. I would also like to thank Dr.
Nosoudi for helping and training me during the research. I have gratefully appreciated all
the assistance she provided through giving me her time and her efforts. Also, I would like
to thank my thesis committee member Dr. Sunar. Also, I would like to thank my colleague
lab students.
In addition, I would like to thank my husband, Ali for making me follow my dream, for
encouraging me, and for providing me the suitable atmosphere to focus on my study
without him I could not be here. Also, I want to thank my mother and my father for their
love and support.
1
CHAPTER I
Introduction
1.1 Project overview
Bone fractures represent a substantial incidence and cost burden among
musculoskeletal injuries in the United States, with 15.3 million cases reported each year,
of which 5-10% involve complications, such as delayed or non-unions (Nauth, et al., 2011).
Conventional surgical treatments for bone damaged that resulted from trauma, tumor, bone
fracture and abnormalities when the defect is above the critical size of 1-mm by either
using autografts, allografts or metallic or ceramic implants (Meng et al., 2010). Each of
these bone grafts has some limitations, such as donor site morbidity, the risk of infections,
pain, shortage in graft quantity and pathogen transmission (Brydone, et al., 2010; Meng, et
al., 2010). Thus, recently, extensive efforts have been done to enhance bone tissue
engineering to mimic the bone tissue microenvironment focusing on scaffolds as
substitutes autologous or allogeneic bone grafts (Xu, et al., 2014)
Bone tissue engineering and regenerative medicine required three complementary and
essential components to mimic the native tissue through culturing diverse cell types in
three-dimensional (3D) microenvironments to generate a fully functional tissue. These
components are osteoprogenitor cells that responsible for creating bone tissue, biomimetic
scaffold to provide 3D structure that is osteoconductive, and stimulation which can be
biochemical or biophysical in nature, to create an osteoinductive environment for
2
promoting cell differentiation (Horii, et al., 2007; Stevens, 2008; Neves, et al., 2017). To
generate this microenvironment requires that the scaffold is manufactured in a way that
mimics the nanostructure of the natural extracellular matrix (ECM; Li, et al., 2013). The
ECM is considered a vital component of any tissue, which works not only supporting cells
but also by creating an appropriate microenvironment that influences cell-function (Horii,
et al., 2007). Therefore, by using a biomimetic scaffold that has similar properties as
natural tissue ECM seeded with suitable cell types under the presence of appropriate
stimulation can guide tissue growth and regeneration.
However, creating living tissue constructs that are structurally, functionally and
mechanically comparable to the natural bone has been a challenge (Polo-Corrales, et al.,
2014). For instance, recreating the bone tissue microenvironment using the appropriate
combination of cells, scaffold and stimulation to direct differentiation (Polo-Corrales, et
al., 2018). As an alternative, the use of extremely-low-frequency pulsed electromagnetic
fields (pEMFs) as an adjuvant therapy for the treatment of bone disorders to reduce
complications have been widely used in orthopedics since it was approved by the FDA
almost four decades ago (Lohmann, et al., 2000). This allowed various clinical trials and
production of commercial devices to promote bone fracture healing (Heckman, et al.,
1981). Since then, different effects of pEMF stimulation on in vitro differentiation and
proliferation of osteogenic cell lines have been published in the literature (Daish, et al.,
2018). Researchers have indicated that the forced vibration of all the free ions on the
surface of a cellular plasma membrane, changes in voltage, and conductivities are possible
mechanisms in which EMF could regulate cell process (Panagopoulos, et al., 2002;
Markov, 2007; Ross, et al., 2015; Rubio Ayala, et al., 2018). Since then, man investigations
3
have focused on the use of this therapy to accelerate the cell proliferation and osteogenic
differentiation of progenitor cells, such as mesenchymal stem cells (MSCs).
A strategy to improve fracture healing is to look more closely at the bone tissue
microenvironment, in which mechanical stimulation is an important part and is essential
for bone health and homeostasis (Voog and Jones, 2010). The importance of this type of
stimulation is because it mediates an adaptive remodeling response in the bone at the
cellular level, through a process known as Wolff’s law (Lanyon, 1974; Lanyon and
Baggott, 1976; Woo, et al., 1981). At the molecular level, the process by which cells
transduce these force-induced signals into biochemical responses is known as
mechanotransduction and it leads to variations in gene expression, cell function,
morphology, and extracellular matrix (ECM) remodeling. Since the mechanisms involved
during bone remodeling are the same as those found during fracture healing (McKibbin,
1978), it is thus reasonable to consider mechanical stimulation as a therapeutic strategy to
induce healing when a bone fracture is present. A recent alternative that has shown great
interest is the use of scaffolds that can provide mechanical stimulation through vibrations
induced on superparamagnetic scaffolds by external EMF (Zeng, et al., 2012; Meng, et al.,
2013; Xu and Gu, 2014; Grant, et al., 2015). In addition, the presence of superparamagnetic
nanoparticles (NPs) could also affect the mechanical properties of the scaffolds, by
enhancing the compressive strength and elastic modulus. The improved mechanical
properties in 3D scaffolds, particularly elastic modulus values, have been proven to
promote MSC osteogenic differentiation (Jegal, et al. 2011). Early work on this type of
scaffold was focused on polymer-based nanocomposites loaded with magnetic particles for
drug delivery applications (Edelman, et al., 1984; Liu, et al., 2006; Zhao, et al., 2011).
4
Recently, the use of these superparamagnetic scaffolds and EMFs, to promote bone
formation has been proposed (Kanczler, et al., 2010; Sapir-Lekhovitser, et al., 2016). The
interesting aspect of this approach is that it couples two separate stimuli on the cells,
mechanical, through the vibratory movement of the scaffold where the cells are attached
to, and magnetic, where the forced-vibration of all the free ions on the surface of the plasma
membrane, changes in voltage, and conductivities affecting cell function (Panagopoulos,
et al., 2002; Deng, et al., 2007; Garner, et al., 2007).
In this research, we used a biomimetic scaffold made of the self-assembling peptide
RADA16, which consist of regular repeats of alternating ionic hydrophilic and
hydrophobic amino acids and self-assemble to form stable β-sheet structures in water.
When exposed to physiological solutions they spontaneous assemble into a stable,
macroscopic membranous matrix, composed of ordered filaments (~10 nm) forming pores
5–200 nm in size (Wang, et al., 2008; Zhang, et al., 1993; Zhang, et al., 1995). the self-
assembled peptide is the spontaneous arrangements of amino acids that followed the state
of thermodynamic equilibrium to get a well-defined structure with firm organizations
through various noncovalent interactions to produce hierarchical structures. In spite the
fact that these interactions are weak, but when they together it formed a firm and stable
structure (Zhang, 2002). This type consists of an alternating amino acid (Arginine, Alanine,
and Aspartic acid) that spontaneously assembling to produce microscopic and macroscopic
matrices of interwoven nanofibers to form higher-order hydrogels scaffolds in the presence
of monovalent cations (Zhang et al.,1995; Hauser et al., 2010; Zhao et al., 2006). Although
these interactions are weak, together they form microscopic and macroscopic matrices of
interwoven nanofibers to form higher-order hydrogels that are firm and stable (Zhang,
5
2002; Hauser et al., 2010; Zhao et al., 2006). Due to their nanometer scale, their network
and biomechanical properties are comparable to the natural ECM, making them good
candidates to generate biomimetic cell niches (Semino, 2008). In addition, their stiffness
can be modulated with concentration. Studies have found that collagen I and RADA16
hydrogels have similar mechanical properties with mean G’ modulus values in the 10–1000
Pa range (Semino, 2008; Cunha, et al., 2011). The generation of superparamagnetic
scaffolds involves either a dip-coating method of the scaffolds in aqueous ferrofluids
comprising biocompatible and nontoxic superparamagnetic nanoparticles (NPs), allowing
these to infiltrate to the scaffold, or using in situ method by mixing NPs during scaffold
synthesis, reducing the number of processing steps and time. It is expected that the
superparamagnetic NPs will chelate to the hydrogel matrix (Kantipuly, et al., 1990). In this
manner, the resulting magnetic scaffolds are capable to take up seeded cells (Bañobre-
López, et al., 2011; Samba Sivudu, et al., 2009). Our focus on this study is on the use
adipose-derived mesenchymal stem cells (hASCs) encapsulated with self-assembled
peptides containing superparamagnetic iron oxide NPs and study their capacity to
differentiate after being biochemically-stimulated with osteogenic induction media and/or
extremely low-frequency pEMFs.
1.2 Objectives
The main goal of this research is to create biologically compatible superparamagnetic
scaffolds suitable for hASC growth and differentiation by mimicking the bone
microenvironment to provide a level of osteoconductive and osteoinductive properties
equivalent native bone tissue. These goals will be achieved through:
6
• Synthesis of biomimetic hydrogels using based on self-assembled peptides
combined with superparamagnetic iron oxide nanoparticles to create a 3D porous
scaffold and evaluate its specific functional groups through FTIR analysis.
• Encapsulate the hASCs within the scaffold and evaluate their viability and
morphology within the scaffold through LDH analysis and F-actin staining.
• Evaluate the hASCs differentiation under pEMF stimulation in vitro by monitoring
the ALP activity and calcium deposition.
7
CHAPTER II
Literature Review
2.1 Bone structure
Bone is a vascularized, metabolically dynamic connective tissue that forms, along
with cartilage, the skeletal system of the body. The main function of bone is to provide
mechanical support and physiological protection. The mechanical support is necessary to
provide protection to vital organs and offer a site for muscle attachment to control
movement. While the physiological protection results from the providing of reservoirs for
growth factors and minerals such as calcium, potassium, carbonate, magnesium, strontium,
chloride and fluoride, phosphate to maintain blood hemostasis and a place for
hematopoiesis to occur. Since bone is a heterogenic structure, these functions vary
depending on its location. For example, the bone that is not exposed to loadings, such as in
the skull and scapula have a dissimilar structure to long bones, which are exposed to applied
external forces such as tension, compression, and bending (Rogel, et al., 2008).
2.2 Bone anatomy
The adult human skeleton, which consists of 213 bones, can be classified according
to their location, shape, or size. In terms of shape, bones can be divided into flat bone
8
(i.e., skull, scapula, sternum, ribs, or mainly the facial skeleton bones) and tubular bone.
The latter can be further classified as long tubular bone (i.e., femur, tibia, humerus,
clavicles, etc.), or short tubular bone as in the hand and feet (i.e., metacarpals, metatarsals,
and phalanges) (Buck, et al., 2012). The anatomical structure of the tubular bone consists
mainly of three different sections: diaphysis, metaphysis, and epiphysis. The diaphysis is
located in the middle of the tubular bone and has the shape of the cylindrical hollow shaft,
composed mainly of dense and hard cortical bone, which is filled by the hollow shaft, or
medullary cavity, that contains bone marrow and fat. The metaphysis located between
diaphysis and epiphysis, is cone-shaped and composed of a growth plate that calcifies with
age. The epiphysis located at the end of the tubular bone, with rounded and wide sections,
consists primarily a trabecular spongy meshwork filled with red marrow and covered from
the exterior with a thin shell of the cortical bone (Clarke, et al., 2008; Buck, et al., 2012).
In contrast, flat bones have a varied structure consisting of either purely cortical bone or
cortical bone with a thin central trabecular region.
At the macrostructure level, cortical and trabecular are the primary bone structures
that formed from the same matrix composition but vary in their porosity, 3D structure, and
their metabolic activities. The cortical (also known as compact bone) makes up 80% of
human body skeletal system, with a porosity of 5-10%, that has a compressive strength
with the ability to resist torsion and bending. Although the trabecular (also known as
cancellous bone) makes up only 20% of the body, with a 50-90% porosity of interconnected
pores filled with bone marrow, which gives the ability of deformation and force absorption.
Nevertheless, both bone structure shares the same metabolic activities since the trabecular
9
bone has a higher metabolic activity (Buck, et al., 2012). In contrast, the cortical bone is
covered by two surfaces, the outer surface called periosteal, which plays a role in the
appositional growth and fracture healing. promoting bone enlargement through the process
of periosteal apposition. The inner surface, called endosteal surface, has a high remodeling
activity, allowing the bone to be under resorption (Clarke, et al., 2008; Datta et al., 2008).
At the microstructure level, the main structure in cortical and trabecular bone
consists of osteons (Figure 1). The osteons (also known as Haversian systems) in cortical
bones are oriented longitudinally along long bones and consist of mineralized collagen
fibers oriented in concentric layers. The osteons are approximately 200 µm in diameter and
10-20 µm in length, with concentric lamellae 3–7 µm in diameter. It is also composed of a
central Haversian canal which is a hollow tube, 80 µm in diameter, located in the center of
the lamellae, allowing blood vessels to pass through it to distribute the necessary nutrients
to bone cells (Athanasiou, et al., 2000; Rho, et al., 1998; Sharir, et al., 2008). Mineralized
collagen fibers that are in a disorganized pattern, known as woven bone, yield a
mechanically weak structure. When the lamellae are oriented tangentially to the external
surface of the bone without making osteons and along with woven bone, they form a larger
plywood-type stacking of dense layer known as a lamellar bone (Rho, et al., 1998).
On the other hand, the trabecular osteons (also known as packets) have a semilunar
shape formed from interconnected plates and rods with a thickness of 50-400 µm and with
300-1500 µm of space between them (Athanasiou, et al., 2000; Clarke, et al., 2008). The
trabeculae distribution reflects the bone strength, shown as an alignment along lines of
10
stress. Consequently, trabecular bone is capable of enduring compressive stresses, an as a
result, it shows a predominant presence in the vertebrae (Datta, et al., 2008).
Figure 1. Three-dimensional structure of bone that shows the cortical and trabecular
components (Buck, et al., 2012).
11
2.3 Bone composition
2.3.1 Bone Extracellular Matrix
Bone is a compound that contains different materials that together constitute the
bone matrix, which consists of about 90% of the total bone. The major components found
in this 3D ECM are about 70% of inorganic components or minerals, 22% of organic
components, mainly collagen, and about 8% lipids and water (Augat, et al., 2006; Buck, et
al., 2012). The distribution of materials and their quality has a strong effect on bone
strength; the composition of the mineral, which is mostly in the form of hydroxyapatite,
gives the bone the ability to resist the compressive load and its quality affects bone
stiffness. The collagen fibrils give the bone the ability to resist the tensile load and its
quality affects bone toughness (Augat, et al., 2006; Athanasiou, et al., 2000). The
mechanical properties of bone result from the interactions between these components, since
the deposited mineral crystal and collagen are oriented in the longitudinal axis, giving bone
its high stiffness and strength along its axis (Athanasiou, et al., 2000). The organic and
inorganic components of the matrix interact together to produce a diverse set of mechanical
and biological characteristics than those produce separately. Hence, the composition of the
matrix plays a vital role in the calcified bone tissue whereas the elasticity is associated to
the mineral phase of the bone and plasticity to the organic components of the matrix
(Landis,1995).
The organic component of mature bone that is considered its main building block.
It consists mainly of the fibrous protein type I collagen that represents 90% of ECM
proteins, which is organized in a fibrillar structure and sometimes onto which, the mineral
crystalizes through gaps between the collagen fibrils. The collagen fibril structure consists
12
of triple helical chains twisted around each other, consisting of two α1 polypeptide chains
and one α2 chain. It structures composed of a sequence of three different amino acids
glycine, proline and hydroxyproline that define the helix-forming repeated motif, which
produces a rod shape molecule 1.4 nm wide and 300 nm long (Weiner, et al., 1998; Rogel,
et al., 2008). During collagen synthesis, preprocollagen (i.e., a single chain of the
polypeptide) becomes procollagen (i.e., a triple helix chain of the polypeptide) after
hydroxylation and glycosylation, where hydrogen and disulfide bonds are formed. In
addition, another bond is generated when tropocollagen (i.e., procollagen with a cut in the
extra terminal amino acids) during which covalent bonds are formed between the terminal
of the triple polypeptide helices (Last, et al., 1984; Zioupos, et al., 1999). The collagen
fibril results from the self-assembling of tropocollagen by establishing a link to
neighboring tropocollagen molecules through trivalent bonds known as
hydroxypyridinium bonds (Zioupos, et al., 1999). The attachment between the NH2-
terminus of one triple helical molecule and the COOH-terminus of the following molecules
forms a gap or a hole. This arrangement forms an area with densely packed molecules and
an area of less densely packed molecules in the 2D model arrangement, while the 3D
arrangement shows transverse channels or grooves (Weiner, et al., 1998). The mineralized
collagen fibrils diameter is approximately 80-100 nm, and the mineral crystals between
these fibrils are around 50 nm by 25 nm and 2–3 nm in thickness (Sharir, et al., 2008). In
contrast, noncollagenous proteins, which represent 10-15% of total bone protein, such as
serum albumin and α2-HS-glycoprotein, have an impact on ECM mineralization through
binding to the hydroxyapatite crystal through their acidic properties. These noncollagenous
proteins can be divided into proteoglycans and glycosylated proteins, which affect the bone
13
mineral deposition and bone cell activity. Of these, osteonectin is considered the most
widespread noncollagenous protein, representing around 2% of the entire protein in bone.
The main impact of this protein on the bone is through its effect on osteoblast proliferation
and matrix mineralization (Clarke, et al., 2008).
The second major component of the bone matrix is bone mineral known as dahllite,
which has a hexagonal crystallographic symmetry, although the microscopic results show
it does not exhibit this type of symmetry instead it shows as a thin plate-shaped crystal
(Weiner et al., 1998). The main mineral component of bone that gives it its hardness
resulted from the presence of mineralized calcium phosphate in form of hydroxyapatite,
[Ca₁₀(PO4)6(OH)2], with small amounts of carbonate, magnesium, and acid phosphate
(Datta, et al., 2008; Weiner, et al., 1998). The mineral crystal has a dimension of tens of
nanometers in length and several nanometers in width, with an elongated morphology and
a preferred crystallographic and morphological alignment with the main directions of stress
(Rogel, et al., 2008). The orientation of minerals has an influence on mechanical properties
of bones, where the longitudinal alignment of hydroxyapatite crystals promotes transverse
isotropy (Sasaki, et al.,1989). In addition, the presence of alkaline phosphatase and other
non-collagenous proteins, such as osteocalcin, osteopontin, and bone sialoprotein help in
ECM maturation. Consequently, these calcium- and phosphate-binding proteins control the
quantity and size of hydroxyapatite crystals, thus influencing the mineral deposition
(Clarke, et al., 2008).
The third main component of bone is water, which has a major influence on bone
mechanical properties as it was found that the dried bone has different mechanical
14
properties than it is wet. The water exists mainly in the holes between the triple-helical
polypeptide and within the fibril and between collagen fibers (Weiner, et al., 1998).
2.3.2 bone cells
Cells in bone tissue makeup 10% of its volume and are composed of different types
that are at different stages of maturity (Figure 2). These cells migrate from different places,
they can have a hematopoietic origin such as the osteoclast or be derived from local
mesenchymal cells known as osteoprogenitor cells such as osteoblasts, lining cells, and
osteocytes (Buck, et al., 2012).
Osteoblasts are mononuclear cells with diameter range between 15 to 30 µm that
has a large spherical nucleus with high composition of rough endoplasmic reticulum and
Golgi apparatus, which promotes ECM synthesis, mainly collagen type I. The presence of
actin, myosin, and other cytoskeletal proteins allows these cells to modify their shape and
helps them during migration and binding to the ECM (Jayakumar, et al., 2010). Osteoblasts
play a vital role in skeletal development by becoming highly specialized synthetic cells,
that when mature, support and regulate hematopoiesis. Besides, they have the capacity to
respond to many mechanical and systemic stimulations, which induce mineral deposition
(Taichman, 2005).
Osteoblasts cover the surfaces of bone and are closely aligned with each other.
Since they are very active, their shape is oval or polyhedral. In addition, during new matrix
secretion, osteoblasts will consist large amounts of endoplasmic reticulum, Golgi apparatus
(for synthesizing proteins), and mitochondria. Thus, osteoblasts not only form and secrete
the organic matrix of bone but also have a role in its mineralization through the control of
15
electrolyte fluxes between the extracellular fluid and the osseous fluid (Buckwalter, et al.,
1995). After becoming activated, these cells can either remain inactive osteoblasts,
osteocytes, or return to osteoprogenitor cells (Buck, et al., 2012). Osteoblasts main protein
secretion is collagen type I, which self-assembles into fibrils. Osteocytes are osteoblasts
that have surrounded themselves with an organic matrix to live within vacuoles called
lacunae. They develop cytoplasmic projections that move across bone to come in contact
with adjacent osteocytes, allowing direct communication (Marks, et al., 1988).
Figure 2. Topographic relationships of different bone cells that initiate from different
origins (Marks, et al., 1988).
Osteoclasts are the only bone cells responsible for the resorption of bone mineral
and originate from CD34+ hematopoietic progenitor cells. They are located on the surface
of bone close to the vascular channels and can be distinguished from the other bone cells
16
by its large and multinucleated cells shape (Taichman, 2005). Once activated, osteoclasts
form specialized membrane structures, such as the sealing zone, where they attach to bone
and the ruffled border, through which they release hydrogen ions to aid in the mineralized
bone matrix dissolution. Cytoplasmic vacuoles, containing primary and secondary
lysosomes can also be seen close to the ruffled border area (Marks, et al., 1988).
Osteocytes are the most abundant type of cells in bone tissue, making up to 95% of
bone cells, being ten times more common than osteoblasts. In spite the fact that osteocytes
are buried in the bone mineralized matrix, they can still communicate and connect with
each other and with cells located on the bone surface using a network of cell processes that
cross their path through canaliculi in the ECM (Franz‐Odendaal, et al.,2006). These
osteocytes are in the interior of the bone tissue and reside within spaces known as lacunae.
In addition, these cells have the capability to make and resorb bone to change the volume
of its lacunae (Marks et al., 1988).
Bone lining cells, also known as resting osteoblasts, cover the surface of the bone
to provide a selective barrier between bone and other extracellular fluid compartments.
Besides, they represent one of the two states of terminal differentiation of osteoblasts.
These cells have flat elongated morphology separated from the marrow (Marks, et al.,
1988). These cells do not achieve any bone formation or resorption, and they act as
gatekeepers, informing if the bone needs remodeling. They can receive and deliver signals
to other bone cells, as well as to those in nearby soft tissue (Parfitt,1994). Since bone lining
cells are inactive, they require fewer organelles and cytoplasm than osteoblasts. Their
function still not clear, although some reports have shown they secrete enzymes to dissolve
17
the osteoid matrix on the bone surface, allowing osteoclasts to remove the bone in the
presence of parathyroid hormone (Downey, et al., 2006).
2.3.3 Growth factors
Growth factors and cytokines provide biochemical stimulation to cells within bone
tissue inducing bone growth and repair. They are soluble polypeptides that target specific
cells, binding them through transmembrane receptors (Vo, et al., 2012). There are many
growth factors with the capacity to affect bone cell function and provide the necessary
induction for a bone formation, also known as osteoinduction. The way growth factors
induce changes in cell function is through something called signal transduction and
involves protein phosphorylation, ion fluxes, changes in metabolism, gene expression, and
protein synthesis. They are different from the hormones in the way of delivery and response
(Lee, et al., 2011). Growths factors found in bone tissue, such as, bone morphogenic
proteins (BMP), insulin-like growth factors (IGFs), fibroblast growth factor (FGF),
transforming growth factor β (TGF-β), and platelet-derived growth factor (PDGF), affect
both cartilage and bone cells (Wozney, et al., 1988).
One of these growth factors BMPs have over sixteen subtypes classified due to their
structure similarity (Haarman, et al., 2005). These originate from the TGF-β superfamily
and play an important role in bone growth during embryonic development and during adult
bone healing. Therefore, the studied elucidates that BMP-2 through 7 and BMP-9 are the
most widely studied osteogenic molecules, which can provide the required signal for the
differentiation of MSCs cells into osteoblasts (Harman, et al., 2005). During bone healing,
BMPs molecules can stimulate progenitor cells in the bone-formation cascade to promote
18
healing of the damaged tissue (Calori, et al., 2009). BMPs are now available commercially
in a form of recombinant human (rh)BMP-2 and rhBMP-7, which have been approved for
restricted clinical use, such as, for non-union, bone defects, open tibial fractures, and spinal
fusions (Nauth, et al., 2011).
It has been established that various members of the TGF-β superfamily are involved
in many functions related to bone induction, such as embryonic growth, tissue
morphogenesis, cell proliferation and differentiation (Vo, et al., 2012). Also, IGF-I and
IGF-II are known to be involved in maintaining osteoblast function and enhancing cell
proliferation and differentiation through the stimulation of precursor cells and the
upregulation type I collagen expression. Besides, studies detected that downregulation of
IGF-I expression leads to apoptosis of bone cells (Canalis, 2009)
Another growth factor is known to play a vital role in the proliferation of
mesenchymal stem cells (MSCs) and preosteoblast differentiation to osteoblasts is PDGF
(Khojasteh, et al., 2013). This growth factor consists of a disulfide-bonded dimer of two
homologous polypeptide chains, which acts as stimulating factor for the stimulation of
angiogenesis, which is an essential for bone healing (Elangovan, et al., 2014).
The FGF is an extremely potent mitogen for mesodermal cells by initiating
angiogenesis required for wound healing; hence, it is important in bone fracture healing by
providing an early vascular response through the formation of new microvessels (Kigami,
et al., 2014). FGF-1 (acidic FGF) and FGF-2 (basic FGF) helped to promote angiogenesis
and osteoblast proliferation, enhancing fracture healing (Sela, et al., 2012).
19
2.4 Bone diseases
Bone tissue might get damaged either by disease, fail while bearing a mechanical
load, or from hormonal deficiencies that make the bone tissue predisposed to damage.
Bone defects include fractures, osteoporosis, osteogenesis imperfecta, osteomalacia,
cancer, and Paget's disease.
Fractures result from either trauma or excessive mechanical strain that leads to
discontinuity of bone tissue (Velasco, et al., 2015). Osteoporosis is an age-related disorder
of the bone tissue microarchitecture and results from a loss in bone density. There are
multiple causes leading to osteoporosis, for instance, hormone deficiency, poor nutrition
from calcium and vitamin D deficiency, sedentary physical activity, and several
pharmacological agents (Downey, et al.,2006). Osteogenesis imperfecta is a genetic
disease that makes the bone brittle and fragile, due to a collagen matrix disorder caused by
a mutation in one of the two genes that encode the chains of collagen type 1 (i.e., COL1A1
and COL1A2) (Rauch, et al., 2004). Osteomalacia is disease resulting from nutritional
deficiencies that allow the bone to lose mineral (Velasco, et al., 2015). Cancer of bone
tissue causes a massive pain and bone destruction. Metastatic cancer causes various effects
on bone, such as an increase in calcium ions in the blood, the release of inflammatory
mediators, which leads to the disintegration of the cortical and trabecular bone,
predisposing to bone fractures (O’Toole, et al., 2006). Paget’s disease is a common aging
disease affecting bone by inducing a high rate of bone remodeling, which results in
abnormal bone formation. The cause of this disease can be either genetic or environmental,
20
and the main sites that commonly affected are the pelvis, vertebrae, and the femur (Al
Nofal, et al., 2015).
2.5 Bone replacement and tissue engineering
Due to increasing the number of patients with bone disease in the United States,
with osteoporosis affecting over 30% of the population have led to an annual cost burden
of $200 B. Orthopedic surgeons estimate that 50% of women with age exceeding 65 years
are at high risk to develop bone fractures (Luu, et al., 2009). The cases where the bone
defects are larger than the critical size of 1 mm, in which they can heal without assistance,
are of specific concern (Hollinger et al., 1990). The current treatment for such types of
defects involves the use of bone grafts, mainly autografts, allografts, and synthetic grafts
(ceramics and metals). The use of these bone graft is estimated at around 600,000
procedures annually in the United States, and approximately 2.2 million worldwide (Panek,
et al., 2015).
The autograft is a bone graft used to restore integrity through the transplantation of
healthy bone tissue from the same patient into the defect site, which helps mitigate immune
rejection issues. This type of graft is considered ideal since it incorporates both the
osteogenic cells and the mineralized matrix from the individual. The most common place
used to harvest the bone from is the iliac crest. However, autografts have some limitations
which limit its use, for example, donor site morbidity, pain, bleeding, infection, the struggle
in shaping the bone grafts to fill the defect, and it is not suitable for large bone defects since
it needs a large bone mass (Brydone, et al., 2010; Rose, et al., 2002).
21
The allograft is a bone graft harvested from a donor usually a cadaver which is
processed and preserved for its future use. Although the allograft has more flexibility to be
shaped according to the defect site, it raises the potential problems, such as immune
rejection and pathogen transmission (Salgado, et al., 2004; Rose, et al., 2002).
Nevertheless, some procedures sterilize the bone graft, which has a negative influence on
the osteoinductive properties required for bone healing, leading to longer healing time than
autografts (Parikh, 2002).
Another substitute used to treat large bone defects are synthetic grafts such as those
made from metals and ceramics. Both grafts work to provide the necessary mechanical
support, but they have limited healing performance. For instance, metal grafts have poor
integration with the host tissue at the injury site, which can lead to infection. Ceramics have
limited tolerance to torsion, bending, or shear stress due to its low tensile strength and are
brittle (Salgado, et al., 2004).
Still, there still an unmet need of developing an ideal bone replacement, which will
require further work in bone tissue engineering and regenerative medicine, to provide the
ideal substitute for treating bone defects. Tissue engineering is a developing field focusing
on repair, replace, or regenerate tissue by combing physics, chemistry, and biology to
create effective materials (Griffin, et al., 2004). The objective of tissue engineering and
regenerative medicine is to restore, maintain, enhance tissue function to make in vitro tissue
to implanted later in vivo; this goal can be achieved by the presence of three main
components, mainly, a biocompatible scaffold, a cell source, and growth/differentiation
conducting cell culture conditions (Kannarkat, et al., 2010).
22
2.6 Cellular sources for bone tissue engineering
For optimal bone regeneration and tissue engineering applications, it is important
to use a reliable cell source to ensure no immune rejection, with osteogenic potential, a
strong proliferation rate, and that allows good integration with the host tissues while
offering the ability for load-bearing and remodeling (Logeart-Avramoglou, et al., 2005).
The most common types of stem cells that used for bone tissue regeneration are MSCs,
embryonic stem cells (ESTs), and induced pluripotent stem cells (iPSCs). The main
sources of MSCs are bone marrow (BM-MSCs), and adipose tissue (ASCs), which are
usually cultured within 3D scaffolds to generate a new bone tissue with osteoinductive cues
(Griffith, et al., 2004).
The most common type of cells used for bone regeneration is adult stem cells, such
as MSCs, due to their capacity to self-renew and are readily available. Also, they can be
broadly proliferated in vitro, besides under specific culture conditions, they can
differentiate into various phenotypes, such as bone, cartilage, muscle, marrow stroma,
tendon/ligament, fat, and different connective tissues (Campagnoli, et al., 2001; Caplan,
2005; Vo, et al., 2012). In addition, in vivo behavior makes them an excellent candidate for
tissue engineering applications due to their regenerative response to injury or disease
(Caplan, 2005). Stem cells generate progenitor bone cells, which during development
become pre-osteoblasts, and osteoblasts (Griffith, et al., 2004). Stem cells are
undifferentiated cells which can become specialized cells and tissues in response to specific
stimulation, which can be biochemical and biophysical in nature. However, the life cycle
of stem cells from totipotent stem cells (morula stage), capable of differentiating into all
23
embryonic and extraembryonic tissues. Then, their potency reduces to pluripotent stem
cells (blastocyst stage), creating all embryonic tissues and to multi- or unipotent adult stem
cells, forming tissues within their specific germ layer (Panek, et al., 2015). MSCs are
culture-adherent, multipotent progenitor cells, and can be isolated from numerous tissues,
for example, bone marrow, adipose tissue, muscle, amniotic fluid, human placenta,
periosteum, cord blood and even peripheral blood (Vo, et al., 2012).
Pluripotent human embryonic stem cells (hESCs) form during embryonic
development, isolated from the inner cell mass of blastocysts, have the capacity of
differentiating into any type of specialized cell found in the bone. Some studies
demonstrate that hESC-derived mesenchymal progenitors could have the same cellular
behavior as the adult BM-MSCs (Marolt, et al., 2012). Since they are pluripotent, they can
differentiate into cell types from all three germ layers. However, hESC raised ethical
concerns, which greatly limited their clinical use since it required sacrificing embryos
(Bitar, et al., 2014; Vo, et al., 2012). Some studies showed that using hESCs might lead to
the formation of a teratoma and that it might also cause immune rejection (Liu, et al., 2014).
Induced pluripotent stem cells (iPSCs) are artificially made pluripotent by making
an adjustment in the expression of specific genes. Some studies have shown that human
iPSCs exhibit some featured like hESCs, for instance, their morphology, gene expression,
surface antigens, and in vitro differentiation ability and pluripotency. Thus, making them
a viable alternative to hESCs for tissue engineering, and thereby eliminating any restrictive
issues or ethical concerns (Teng, et al., 2014). Nevertheless, it is still important to continue
with further research to define what is the best starting somatic cell for iPSC generation for
24
human clinical research (Amini, et al., 2012). Until now, most of the bone tissue
engineering methods developed are based on the use of easily-accessible BM-MSCs as a
cellular source. These have shown the ability to differentiate into chondrocytes and
osteoblasts in vitro (Colnot, 2011). BM-MSCs are currently considered the choice for bone
tissue engineering and regeneration applications due to their high osteogenic ability
(Yousefi, et al., 2016). However, the BM-MSCs have been shown to decay with growing
passages, requiring longer periods to proliferate, thus showing limited proliferative
capacity (Bruder, et al., 1997).
The stem cells used in this research are human adipose tissue (hASCs) since these
very accessible, requiring only local anesthesia with slight patient discomfort, through
procedures like, which is less invasive than bone marrow aspiration. Also, it found that 1g
of adipose tissue contains about 5×103 hASCs, which is 500-fold higher than in bone
marrow (Yousefi, et al., 2016). In addition, different studies imply that hASCs can commit
to osteogenesis quickly and with less biochemical stimulation (i.e., through exogenous
cytokines), making them better candidates for bone tissue engineering (Levi, et al., 2011),
with equivalent potential to BM-MSCs of differentiating into cells from the mesodermal
layer (Mizuno, 2009). Moreover, increasing evidence suggests that hASCs and BM-MSCs
are distinct cell populations that differ in their inherent properties (Macotela, et al., 2012;
Tchkonia, et al., 2005; 2006). This is evidence through their surface antigen expression
profiles, with hASCs being STRO-1 negative, CD36 positive and CD106 negative in
contrast to BM-MSCs. In general, different surface markers have been reported to define
hASCs, such as CD44 (hyaluronate) and CD90, as well as integrin b1 (CD29), endoglin
(CD105), and integrin a4 (CD49) (Levi, et al., 2011). In addition, an in vitro study using
25
hASCs showed these could differentiate into osteoblasts as early as 3 days after using
induction media (Wan, et al., 2006).
2.7 Three-dimensional Scaffold for bone tissue engineering
To mimic the native tissue structure, 3D scaffold designed which is a fundamental
tool in tissue engineering able to mimic the bone tissue microenvironment by modifying
its composition, structure, surface chemistry, and stiffness (Marolt, et al., 2012). This
design must control cell response, such as adhesion, migration, and proliferation (Causa,
et al., 2007). Therefore, 3D porous scaffolds must provide surface and void volume that
allows proper cell attachment, migration, proliferation, and differentiation to generate the
desired tissue (Griffith, et al., 2004). Thus, the scaffold should be designed following a
specific criteria appropriate for bone tissue engineering, for instance, it should be
biocompatible and not induce an immune response, be nontoxic, absorbable at a rate
proportional to bone formation, be osteoconductive, capable of being sterilized, and easy
to manufacture and handle (Logeart Avramoglou, et al., 2005; Gunatillake, et al., 2003).
Moreover, the scaffold should promote cellular attachment, provide a porous structure to
allow cells to migrate, differentiate and secrete their ECM, allow cells to integrate with
tissue, and be in contact with bioactive molecules (Caplan, 2005). However, in the past
decades, many efforts have focused on the discovery of nobel scaffolds materials that can
be better substitutes for autologous or allogeneic bone grafts in the bone regeneration (Xu,
et al., 2014). These have been developed by using either synthetic, natural, or composite
26
materials from organic or inorganic sources. These include synthetics such as calcium
phosphates, which are inorganic, while polymers like poly(phosphazenes) poly(tyrosine
carbonates), poly(caprolactones), poly(propylene fumarates), and poly(hydroxy acids) are
organic (Karp, et al., 2003). In spite the fact that using natural materials have many
benefits, synthetics offer benefits many of the former cannot, such as not having the
immunogenic problem of natural biomaterials, or have a much better control over material
properties. However, a major deficiency of synthetics is due to their limited biological
recognition capacity (Causa, et al., 2007). In search of a balance, in this study, we used the
self-assembled peptides, which are synthetic but based on amino acids, which are natural
biomaterials to provide a biomimetic 3D microenvironment for hASCs.
2.7.1 Self-assembled peptide scaffold
The main building block of self-assembly molecular system is chemical
complementarity and structural compatibility connected through weak noncovalent bonds
such as hydrogen bonds, hydrophobic bonds, van der Waals interactions, and ionic
complementary bonds under thermodynamic equilibrium settings. These bonds gather the
molecules into stable organized structures (Zhao, et al.,2006). Self-assembled peptides
have diverse applications extending from models to evaluate protein folding and protein
conformational diseases to their application in tissue engineering and regenerative
medicine as 3D scaffolds that mimic the microenvironment of native tissues and drug
delivery. Due to their capacity to generate a diversity of structures, properties, being easily
synthesized at a reasonable price, allow these self-assembled peptides as candidates in
numerous technological innovations (Zhao, et al.,2006). Also, due to the peptidic nature of
27
these provides a great opportunity for in vivo applications since the degradation products
of this peptides have a lower immune response, when compared to other (Cormier, et al.,
2013).
In 1989, a study was done by Shuguang Zhang on yeast genetics and protein
chemistry led to the discovery of a self-complementary peptide. He recognized a protein
named zuotin with the 16-residue peptide sequence repetitive motif, n-
AEAEAKAKAEAEAKAK-c (EAK16-II), between a segment of interchanging alanine-X
repeating 34 residues. Later, he performed different analysis to evaluate its structure and
its biochemical properties to produce a class of simple β-sheet peptides (Hauser, et al.,
2010). In this context, different types of peptides have been developed from this class, such
as RADA16-I, RAD16-II, EAK-I, and EAK16-II, which all self-assemble into stable β-
sheet structures in aqueous solution to form a nanofiber that can be used as a 3D scaffold
for tissue engineering (Hauser, et al., 2010).
Self-assembled peptides type I are those with alternating hydrophobic and hydrophilic
sequences that connect through different noncovalent bonds, forming a β-sheet structure in
aqueous solution that contains monovalent alkaline cations or in physiological media
(Zhang, et al., 1994). This behavior results from the hydrophobic side chains on one side
of the charged amino acid and hydrophilic side chains on the other surface that enabled it
to be assembled into specific structure as pegs and holes are known as ‘‘molecular Lego’’
according to its appearance at molecular level (Zhang, et al.,1993; Zhang, 2002). This
alternating of amino acid residues results in the formation of β-sheets with a nanofiber with
28
a hydrophobic core and a hydrophilic surface resulting from the stacking of two β-sheets
into a simple fibril unit (Cormier, et al., 2013). Also, some studies showed that combining
different type of polymers into the alternating amphiphilic-peptide, reveal that these
polymers can form β-sheet structures that can aggregate, depending upon pH, salt, and time
(Zhang, et al.,1993). In contrast, peptides that have the same amino acids composition but
with a different sequence, they tend to form α-coils structure instead of stable β-sheets
(Zhang, et al.,1994). These stable self-assembled nanofibers are affected by the pH and
ionic strength of the aqueous solution (Nagai, et al., 2006). Once the type I peptides are
assembled, they remain stable under varying chemical and physical states. Therefore, they
demonstrate resistance to degradation by several proteases, heat, and chemical denaturation
agents (Zhang, et al.,1993; Zhang, et al.,1995).
One member of the type I family is RADA-16 (Figure 3), which is commercially
available as PuraMatrixTM that has been used in this research (Zhao et al.,2006). This type
of self-assembled peptide produces microscopic and macroscopic matrices of interwoven
nanofibers that form higher-order hydrogels in the presence of monovalent cations (Zhang,
et al.,1995). The resulting hydrogel is based on 10-20 nm diameter fibers, 5-200 nm pores
and a 99% water content (peptide content 1-10 mg/ml) (Wang, et al., 2008; Zhang, 2002).
Figure 3. Amino acid sequence of self- assembled peptide (RADA-16) (Zhang, et
al., 1993).
29
The amino acid sequence of this type contains positively charged Arginine,
Alanine, and negatively charged Aspartic acid, forming the RADA16-I peptide (AcN-
RADARADARADARADA-CONH2). The RADA-16 peptide is ionic self-complementary
due to the presence of an ionic pair between the positive amino acid residues (arginine) and
the negative amino acid residues (aspartic acid). Consequently, this sequence of amino
acids leads to the formation of two distinctive hydrophobic and hydrophilic sides, whereas
the hydrophilic sides represent the outside of the peptide fiber that is in direct contact with
water, and the other side form a double sheet inside the peptide fiber (Hauser, et al., 2010).
Due to the presence of both the polar and the non-polar surface, a stable β-sheet is formed,
considered essential for peptide self-assembly and nanofiber formation (Wang, et al.,
2008).
In aqueous solution, the PuraMatrix forms hydrogen bond through its backbone;
besides they contain two distinctive sides, one hydrophobic due to the overlapping of
alanine, as with the spider silk or silk fibroin, with the other side of the backbone being
hydrophilic due to arginine and aspartic acid (Yokoi, et al., 2005). Although the forces
generated during sonication can disassemble the RADA-16 hydrogels, they will readily re-
assemble the moment sonication ceases. Sonication acts to mechanically break the
hydrogen, ionic, and hydrophobic bonds to produce peptide fragments. This result when
the hydrophobic bonds that formed between alanine and water are disrupted mechanically,
during which time their cohesive ends start to find each other by sliding diffusion in
aqueous solution due to it being energetically unfavorable (Yokoi, et al., 2005). These 16-
30
residue peptides have a length between 2.5 and 5 nm, and after reassembling they form a
longer nanofiber between a hundred nanometers to a few micrometers in length (Yokoi, et
al., 2005). Self-assembled peptides are commercially synthesized either through a solid
phase or solution peptide synthesis chemistry (Hauser, et al., 2010). Since this self-
assembled peptide scaffold contain a large amount of water, whereas water molecules able
to be arranged by surface tension to create clusters divided by nanofibers into
compartments. The 3D nanofiber scaffold was able to create an environment that mimics
the in vivo conditions when cells create molecular gradients within the scaffold (Hauser, et
al., 2010).
2.8 Biophysical stimulation
2.8.1 Effect of mechanical stimulation
Bone has the capacity to sense and adapt to any skeletal loading making bone a
dynamic tissue capable of modifying its mass, strength, and geometry to accommodate for
any external mechanical stimulation. This stimulation can affect osteogenic cells by
causing a local deformation of the ECM, generating a fluid flow that causes shear stresses,
and the initiation of electric fields (Mauney, et al., 2004; Zimmerman, et al., 2000). In
general, all eukaryotic cells are sensitive to mechanical and physical forces, such as gravity,
tension, compression, and shear that lead to modulation of cell function. Hence, mechanical
stimulation can initiate a biochemical signal that can be interpreted as a cellular reaction in
a process called mechanotransduction. In this process, mechanical energy is transformed
into electrical and/or biochemical signals (Burger, et al., 1999). Since bone is stiff,
exposure to a physiological load results in a small 0.2% deformation, while in vitro a 1 to
31
3% deformation is required to initiate a cellular response (Burger, et al., 1999). This leaves
shear stress due to hydrostatic pressure and fluid flow, as the main mechanical stimulating
forces (Huang, et al., 2010). Mechanotransduction can be described as the conversion of
mechanical forces to biochemical signals that alter cell function through four different
phases, which are: mechanocoupling, biochemical coupling, transmission of a signal to the
sensory cell, and the effector cell response (Huang, et al., 2010).
To understand the mechanical stimulation from a cellular perspective, the applied
forces on bone cause bone cells to deform and expose them to shear stress from the
interstitial fluid motion in the canalicular spaces. In addition, fluid flow causes streaming
effects that generate electric potentials (Pavalko, et al., 2003). This theory, called
mechanosomes, regards the multiprotein complexes that represent the focal and cell
adhesion protein complexes, the cytoskeleton, the muscleoskeleton and adherents’ junction
protein that connect the neighbor cells and how they respond to load-induced deformations.
This cause changes in ions channels which drive variations in protein conformation
(Figure 4). This leads to the release of protein complexes known as mechanosomes, able
to transfer mechanical information into the nucleus, causing alterations in DNA geometry
and mediating the formation or mobilization of signaling complexes as the mechanical load
transforms into chemical energy (Pavalko, et al., 2003).
32
Figure 4) The mechanosensors that resulted from changes in ions channels,
proteins, and cytoskeleton to activate the intercellular signals (Rubin, et al., 2006)
Therefore, when macroscopic loads generated, osteocytes which, act as the main
sensory cells in bone, senses the interstitial fluid flow generated within the lacuna and the
canaliculi, sending signaling molecules to either osteoclasts for bone resorption or
osteoblasts for bone formation (Pavalko, et al., 2003; Huang, et al., 2010). As fluid flows
within the lacunar-canalicular porosity, load-induced mechanical strains on cells actin
filament of the cytoskeleton which are 1-2 orders of magnitude larger than whole-tissue
level strains, resulting in intracellular signaling (Han, et al., 2004). Hence, this hypothesis
suggests that these strains resulting from canalicular fluid flow act as a local force, instead
of loading-related strains. Thus, when force is applied on the bone, fluid would squeeze
out of the unmineralized matrix adjacent to cell bodies into the Haversian or Volkmann
channels generating fluid shear stress on osteocytes membrane (Burger, et al., 1999).
33
Moreover, canaliculi fluid flow detected on osteocyte surface can generate shear stresses
of 0.8–3 Pa (Burger, et al., 1999). As osteocytes become mechanically stimulated, the start
paracrine signaling to stimulate osteoblasts into the bone formation, including increased
release of nitric oxide (NO) and prostaglandin PGE2 and PGI2, and IGFs (Orr, et al., 2006;
Burger, et al., 1999). The study the mechanotransduction on bone cells have recognized
numerous candidates involved in the mechanosensing process. Mainly, mechanically gated
ion channels, integrins and focal adhesions kinase, G proteins, and the interaction between
the cytoskeleton and certain phospholipase C isoforms. Current researchers have shown
that focal adhesion kinase plays a vital role in mechanically induced bone formation in vivo
(Morgan, et al., 2008).
Since osteoblasts act as an effector cell under proper physical stimulation, these can
initiate osteogenesis. Mechanical stimulators such as stress, strain, and hydrostatic pressure
have been proven to induce bone regeneration and fracture healing (Xu, et al., 2014). In
vitro studies have shown that cyclic pressure improves osteoblast functions related to new
bone development, using a custom-made system that provides cyclically oscillating
pressure with specific amplitude and frequency. These studies demonstrate the
upregulation of osteogenic biomarkers type-I collagen, osteocalcin, and TGFβ1 (Nagatomi,
et al., 2003). Another study using combinations of the two key mechanical stimuli that
affect mesenchymal tissue differentiation (i.e., shear strain and fluid flow), showed bone-
healing through histological analyses (Morgan, et al., 2008). An in vivo study of using 1.2
Pa of fluid flow shear stress for mechanical stimulation revealed an upregulation of COX-
2 and c-Fos expression, which are important to maintain the osteoblast phenotype (Pavalko,
et al., 1998).
34
2.8.2 Effect of electromagnetic fields
Electromagnetic fields play a vital role in tissues development through a cascade of
processes for tissue regeneration that include cell interactions, ECM synthesis, cell
migrations, differentiation, and proliferation. In 1974, electromagnetic fields were
introduced as time-varying fields used for therapeutic purposes, specifically to treat
nonunion bone fractures and congenital pseudarthrosis (Bassett, et al.,1987). After many
studies analyzing the influence of magnetic fields on cells, FDA approved to use pulsed
electromagnetic fields (pEMFs) in 1979 as an active and harmless mechanism for healing
nonunion, congenital pseudarthrosis, and failed fusions (Carpenter, et al., 1994). The idea
behind using EMF to induce osteogenesis come from the natural endogenous streaming
potentials in bone during deformation. Clinical trials using this treatment were carried
through direct conduction by using electrodes; then, they used a wire coil to the general
magnetic field at fracture site through forcing electric currents on it. Later, time-varying
magnetic field was used to generate the required electric field in bone via Faraday coupling
(Pilla, 2002; Funk, et al., 2006). Those devices used pEMFs with extremely low frequency
ranging from 1 to 100 Hz, inducing fields on the microvolt/centimeter level at the fracture
area (Funk, et al., 2006). The reason to use the time-varying magnetic fields is to induce
an electrical field like the one that it generated in dynamically deformed bone tissues with
a similar signal shape and amplitude. The generated fields are used to trigger specific
cellular responses, by relying on the nature of cellular targets, its tissue environment, and
its function (Carpenter, et al., 1994). Recently, pEMFs have been used as a therapeutic
35
procedure to treat pain, inflammation, and dysfunctions related to rheumatoid arthritis and
osteoarthritis (Ganesan, et al., 2009).
FDA-approved noninvasive magnetic fields (pEMFs) used for bone healing and
remodeling, are generated using coils in a Helmholtz arrangement (Figure 5), to initiate
pulses repeated signal with extremely low frequency (15 Hz) and making millivolt per
centimeter (mV/cm) electric fields at the treatment area (Pilla, 2002). Magnetic fields with
a frequency range between 0 Hz up to several hundred GHz are considered as nonionizing
radiation, and researchers have found that this nonionizing electromagnetic energy can
generate different biological effects through interaction mechanisms that do not contain
any macroscopic heating, when the field is applied to tissue samples, it found that the global
temperature change is generally less than 0.001°C (Walleczek, 1991). The magnetic fields
with frequency band between 3 Hz-3 kHz are considered as extremely low EMFs
(Ganesan, et al., 2009).
Figure 5. Schematic illustration of using Helmholtz coil to generate PEMF (Rauh, et al.,
2011).
36
In biological tissue, the effect of EMFs can generate coherently oscillating forces
on charged molecules within the tissue, in a phase with the polarized field and on parallel
planes between each other (Panagopoulos, et al., 2015). Even though the oscillation of the
molecules is at high velocity which results from the thermal motion, it has no biological
effect, the fact that these are coherently polarized allows them to initiate changes at the
cellular level (Panagopoulos, et al., 2015).
To understand the effect of EMFs on tissue engineering and regenerative medicine,
we need to understand the how of electric and magnetic fields interact with the native
tissue. During early embryo development, cells move towards forming an organ, through
migration driven by voltage gradients from by the distribution of charged ions, such as Na+,
Cl-, K+. The currents generated by passive sodium uptake from the environment leads to
an internally positive transepithelial potential difference with an endogenous static electric
field is in the order of 1–5 V/cm. This field is also generated in wounds resulting from
disruption of this transepithelial potential in the epithelial layer.
In bone tissue, the therapeutic use of electric fields is derived from the observation
that when bones are placed under mechanical load (stress) the deformation (strain)
generates an electrical potential. This voltage gradient generates between liquid and solid
from endogenous streaming potentials produced due to fluid motion (Kovacic, et al., 2010;
Funk, et al., 2006). Thus, an EMF is generated due to these processes, which could induce
an electric current in bone tissue by Faraday coupling (Funk et al., 2006). It has been shown
in various studies that there is very close correlation connected between EMF and
mechanical vibration due to the piezoelectric properties of bone (Funk, et al.,2006).
37
At the molecular level, weak EMFs have more influence on cell biology than strong
ones, and its bioeffect is seen in signal-transduction cascades, such as the Ca2+ transport
system, gene expression, cell growth, and apoptosis (Kovacic, et al., 2010). In any typical
tissue, there is intercellular space composed of small narrow fluid channels of 150 Å
between cells that provide conduits for the cell to cell communication. Those channels have
a distinct low impedance as compared to the cell membrane, making them a preferred path
for environmental EMF-induced currents, since it can transmit around 90% of intrinsic
current, leading ionic species to the membrane surface. These channels sense any weak
electrochemical oscillations in the pericellular fluid through the charged tips of
glycoproteins that detect any chemical and electrical signals in the surrounding fluid
(Kovacic, et al., 2010).
There are different theories describing the effect of weak EMFs on cells membrane
and on ion channels. One of these theories is the forced vibration ion theory which states
that there are various kinds of ions surrounding both sides of cell membranes, such as, K+,
Na+, Ca+2, Cl-, etc., which take part in the cell’s signal transduction, and in establishing the
transmembrane electric potential. A flux of ions occurs due to the ion concentration or
electrical potential gradient, which induce their movement through mechanically gated ion
channels. Therefore, when a pulsed low external electric or magnetic field is applied, it will
generate an oscillating force on the cell's membrane, hence on the free ions on both sides
of the membrane and on the ion channel proteins. These vibratory forces will give the
oscillating ions a false signal for gating channels that led to disorder the electrochemical
balance of the cell's membrane and consequently the entire cell function. The ions will be
38
moved with a homogeneous motion where all exhibit the same value and phase. In addition,
it has been suggested that the low-frequency magnetic fields are more bioactive than those
with a higher frequency because there is an inverse relationship between the amplitude of
the forced-vibration and the frequency of the field (Panagopoulos, et al. 2000;
Panagopoulos, et al., 2002).
2.9 Effect of magnetic scaffold under EMF stimulation for bone regeneration
Biophysical forces, mainly mechanical loading and electromagnetic signals are
essential regulators of bone formation. Tagging superparamagnetic NPs to
mechanosensitive cell membrane receptors in osteoprogenitor cells, allows the possibility
of mechanically activating these cells with an external magnetic field enhancing their
osteogenic potential (Kanczler, et al.,2010). Providing the correct stimulatory conditions
that leads to a tissue promoting microenvironment in vitro and in vivo is considered a
crucial goal for regenerative medicine (Sapir-Lekhovitser, et al., 2016). Over the last
decades, superparamagnetic NPs (mainly iron oxide-based) have been widely used in many
biomedical applications (Liu, et al., 2009). When the alternating magnetic fields are
combined with superparamagnetic NPs embedded within 3D scaffold structures, it is
possible to induce mechanical forces on the scaffold to which cells are attached, thus
inducing mechanical forces on the cells themselves (Sapir-Lakhovister, et al., 2016). The
design of novel superparamagnetic scaffolds for bone tissue engineering have generated
much interest in recent years. An approach involves dip-coating of the scaffolds in aqueous
ferrofluids comprising biocompatible and nontoxic superparamagnetic NPs, allowing them
to infiltrate to the pores of the scaffold. In this manner, the resulting superparamagnetic
39
scaffolds are capable to up taking cells and growth factors (Bañobre-López, et al., 2011).
Also, an in situ method can be used to generate a superparamagnetic nanocomposite
scaffold by mixing the superparamagnetic NPs during scaffold synthesis, reducing the
number of processing steps and time (Sivudu, et al., 2009). Most of the superparamagnetic
NPs used are made from iron oxides, such as magnetite (Fe3O4) and maghemite (γFe2O3)
(Dobson, 2008). In bone regeneration, superparamagnetic NPs are widely used for in vitro
as well as in vivo applications, with sizes in the 1–100 nm range. A special characteristic
of interest is their superparamagnetic nature, in which they are only magnetized under the
effect of the external magnetic field, allowing their control noninvasively (Kannarkat, et
al., 2010). Superparamagnetic NP parameters such as size and shape have an impact on
their properties, for instance, their coercivity and magnetization values and the ability to
change the magnetic behavior from the ferromagnetic regime to the superparamagnetic
regime (Jun, et al., 2008). Therefore, by applying a magnetic gradient field to a
superparamagnetic scaffold, causes NP displacement within the scaffold where cells are
bound, inducing compression and tensile forces on the cell membrane, leading to
cytoskeleton deformation and cell dragging. Membrane receptors such as integrins act to
transmit these forces applied on the cytoskeleton through activation of intracellular
signaling pathways to regulate osteocyte and osteoblast function (Russo, et al., 2016).
These superparamagnetic NPs act as magneto-mechanical stimulators of cell arrays at the
cellular level, impacting them through cell proliferation, differentiation, and migration
(Lima, et al., 2015). Therefore, superparamagnetic NPs act as a remote stress inducer
without the need for invasive bioreactor system to mimic the in vivo environment (Dobson
et al., 2006). Furthermore, researchers also utilize superparamagnetic NPs generating
40
magnetic force, using magnetic fields, for inducing drag and rotation. The magnetic drag
technique require NPs coated with a cell-specific ligand that attaches to surface receptors,
such as integrins, to allow vertical displacement when applying magnetic fields. These
forces cause deformation of cell cytoskeleton that activate different mechanosensitive ion
channels. The benefit of this technique is that it produces a localized force on the specific
cell receptor. The other approach is the twisting or rotation method, which works by
magnetizing the NPs in one direction for short pulse then applying a second weaker pulse
perpendicular to the magnetic fields, making the NPs rotate. The advantage of this
approach is that it produced a localized mechanical stress instead of a deformation of the
entire cell membrane (Hughes, et al., 2005). Mechanical stress can also result from
generating a dragging force by magnetic NPs with specific strength and frequency on cell
cytoskeleton, resulting in increased phosphorylation of tyrosine kinases, including the
MAP1 kinase activity (Schmidt, et al.,1998).
Overall, the rate of bone cell growth, proliferation, and differentiation have been
improved by incorporating superparamagnetic NPs into the scaffold in the presence of
external magnetic fields (Kannarkat, et al., 2010). In recent years, nanoscale metals in the
shape of metal oxides in polymer-based nanomaterials have been studied for their enhanced
antimicrobial characteristics in numerous fields (Dhivya, et al., 2015). The research by
Bock and co-workers developed a composite scaffold based on collagen and
hydroxyapatite (HA), the latter was included because of its high osteoinductive properties,
lack of antigenicity or cytotoxicity, and low degradation rate. Collagen was added for its
osteoinductive properties and biocompatibility. Superparamagnetic NPs were added to the
scaffold using the dip-coating method (Bock, et al., 2010). In another study, Kannarkat and
41
his team developed a magnetic scaffold structure that mimics the natural bone tissue and
ECM using polycaprolactone-based scaffolds fabricated by electrospinning. They added
1–100 nm superparamagnetic NPs (Fe3O4) with the polymer precursor prior to
electrospinning. These NPs along with an external magnetic field allow the induction of
low-level mechanical stress within the scaffold, producing shear stresses at the cellular
level. The mechanical stresses on the cells increase the expression levels of multiple genes,
the production of the second messenger nitric oxide and cyclic adenosine monophosphate
and increase in the activity of various proteins. The external magnetic field used was 1–6
Gauss at 15 Hz, applied for two hours daily on MC3T3-E1 mouse preosteoblasts. Their
results showed that the attachment of the cells was similar for all scaffolds with and without
NPs. It was also observed that the cells formed clusters after nine days in culture as signs
of proliferation, but it was slower than the control. They also noticed the cells migrate into
the scaffold due to the high porosity, after cells adhered to the scaffold they exhibited an
elongated shape as a sign of differentiation to osteoblasts (Kannarkat, et al., 2010).
Russo and co-workers used a collagen/HA composite scaffold with
superparamagnetic NPs (mean diameter ~200 nm) for bone regeneration in an in vivo bone
defect model using a rabbit femoral condyle under the effect of static magnetic fields. In
their experiments, they considered four groups, in which the permanent magnet without
superparamagnetic scaffolds was used as controls, while the other two groups fabricated
scaffolds by two methods. The first method synthesized a scaffold made of type I collagen
fibrils-HA mixed with uncoated ferromagnetic iron (II, III) oxide NPs (<50 nm), yielding
freeze-dried cylindrical porous scaffolds. The second method allowed the infiltration of the
NPs into the collagen/HA hybrid porous scaffolds by capillary action. Their results showed
42
that after 12 weeks the bone regeneration noticed with both of magnetic scaffold and the
permanent magnet was significantly higher than with a non-magnetic scaffold and that the
scaffold made using the first method was more osteoinductive than the first Also, the
nanoindentation results showed that the mechanical properties of newly-formed bone
within the first scaffold were closer to that of native tissue and more mature bone trabeculae
as compared to the other groups. In addition, they noticed a clear reorganization of the
scaffold architecture under the static magnetic field in vivo, where the magnetized collagen
fibers aligned in a similar way of the field lines generated by the permanent magnet (Russo
et al., 2016).
In another study, Gloria, et al., (2013) developed a magnetic scaffold for bone tissue
engineering, based on PCL, to which iron-HA NPs were added during polymerization. The
aim was to induce mechanical stimulation on the seeded hMSCs using alternate external
magnetic fields (27 mT, 260 kHz) to control cell function. The magnetic scaffold used three
different polymer-to-particle weight ratios, i.e., 90/10, 80/20 and 70/30 w/w. Results
indicate that the higher the number of NPs embedded in the scaffold the more hydrophilic
and nanostructured would the surface be. They also observed in all groups a magnetically
induced thermal response as a function of NPs during characterization. This is an expected
result due to the high-frequency field used, which can be explained by the energy released
through the Neel relaxation process. Osteogenic differentiation of the hMSCs cells in vitro
by ALP activity showed significant differentiation after a week that the magnetic scaffolds
supporting the osteogenic differentiation. It is important to mention that these magnetic
scaffolds were only exposed to magnetic fields during characterization and not during cell
growth, (Gloria et al., 2013). Also, Panseri and co-workers developed a biomimetic
43
magnetic scaffold, exposed to a rare earth permanent magnet (1.2 T) to assess in vivo bone
regeneration using a rabbit model. Defects 2 mm in size were drilled on the lateral condyle
of the distal femoral epiphysis and filled with one of two different scaffolds, based on
collagen-HA and superparamagnetic iron oxide NPs. Their results showed that the static
magnetic fields have no undesirable impact on tissue formation and these magnetic
scaffolds seemed to have well integrated with an adjacent cancellous bone with no necrosis
or inflammatory response to corrosion products and iron toxicity. After 4 weeks, they
noticed inside scaffolds thin bone trabeculae while in the periphery of the scaffolds were
more mature trabecular tissues. Furthermore, the histological analysis showed a newly
formed woven bone through the scaffolds structure with the presence of normal osteocyte
lacunae developed inside and around magnetic scaffolds, and typical mineralization
gradient initiated from outside to the inside of scaffold. These results confirmed that the
magnetic scaffold along with static magnetic fields has an influence on bone tissue
remodeling and tissue regeneration (Panseri, et al., 2013).
In another study Panseri, et al. (2012) analyzed the effects of the adding different
amounts of magnetic NPs to HA scaffolds (HA/NP 100/0, 95/5, 90/10 and 50/50 wt.%) in
bone tissue regeneration, in an in vitro model using human osteoblast-like cells (Saos-2)
and an in vivo rabbit model under the influence of applying a static magnetic field of 320
mT. Their results showed a significant increase in cell proliferation on the 90/10 scaffold.
ALP activity was measured for all groups, showing no significant increase between the
groups, with no influence by the presence of the magnetic field. On the other hand, the in
vivo experiments showed similar histocompatibility. Also, macroscopic evaluation
displayed no infection or tissue necrosis in none of the scaffolds. Bone tissue was
44
observable around and inside the scaffold in both groups and some pores were full of new
bone, verifying a high level of histocompatibility with the magnetic scaffold comparable
to the control group (Panseri, et al., 2012).
Another in vivo study using superparamagnetic scaffolds combined with the
external static magnetic field using a rabbit model. The scaffolds consisted of a composite
of γ-Fe2O3 and HA NPs in poly(lactide acid) prepared by electrospinning. The magnetic
scaffolds were implanted in New Zealand white rabbits, and to provide the magnetic field
they fixed permanent magnets to the rabbit cages of opposite sides. Their results showed
that at day 10 there were host-derived cells composed mainly of macrophages and
fibroblasts that had migrated to the defect site. Over time, there was degradation of the
implanted scaffold and showed the presence of osteoblast cells and ECM around the
scaffold at day 20 as an indication of a new bone tissue formation. This behavior increased
with time showing an increment in the new bone tissue as the scaffolds degraded. Also, the
superparamagnetic scaffolds exposed to magnetic fields displayed significantly more
collagen than those without magnetic stimulation, which indicates bone growth increased
with magnetic field stimulation. In addition, the new bone tissue became connected and
homogeneous with an organized morphology similar to the original tissue and with faster
degradation rate, while the scaffolds group without external magnetic field exhibited non-
homogeneous bone tissue with slower scaffolds degradation. Thus, the nanofibrous
magnetic scaffolds with the static magnetic field stimulation were able to mimic the
original ECM microenvironment in the defect, promoting osteogenic cell attachment and
growth (Meng, et al., 2013).
45
Zeng, et al., (2012) studied the effect on adhesion, proliferation, and differentiation
of osteoblastic cell lines by superparamagnetic NP content in HA scaffolds after being
stimulated by magnetic fields (1 mT = 10 Oe in Air, 50 Hz). Rat osteoblasts (ROS 17/2.8)
and mice preosteoblasts (MC3T3-E1) were chosen and grown in HA scaffolds with a 70-
80% porosity. with micro- and macro pores with interconnectivity comparable to native
spongy bone tissue, and with different concentrations (0.2 to 2.0 wt.%) of 8 nm
superparamagnetic NPs introduced by dip-coating. Their results showed that the ALP
activity of both cell lines grown in superparamagnetic scaffolds was significantly higher
under magnetic field stimulation than in the scaffolds without NPs. Also, they revealed that
with the increase of NP content within scaffold affects cell adhesion, proliferation, and
differentiation, giving the superparamagnetic scaffold the ability to attain intrinsic
magnetic therapy and gain some synergistic effect to enhance the cell response when
exposed to magnetic fields, allowing them to be used as a bone graft substitutes (Zeng, et
al., 2012).
46
CHAPTER III
Materials and Methods
3.1 Cell culture
Human adipose-derived mesenchymal stem cells (hASCs) purchased from Lonza
(Walkersville, MD) were cultured under standard culture conditions in a sterile, humidified
incubator at 37°C, and 5% CO2/95% air. Cells were cultured in T75 flasks at a density of
5.0×105 cells/flask, using mesenchymal stem cell growth proprietary kit purchased from
ScienCell (Carlsbad, CA) that consisted of 500 ml of basal medium, 5 ml of MSC
osteogenic differentiation supplement and 5 ml of a penicillin/streptomycin solution. The
appropriate amount of growth medium was added to each flask (0.2-0.4 ml/cm2) and was
replaced every three days. The cells used in all experiments were from the fourth passage.
3.2 Three-dimensional cells encapsulation and gel formation
To prepare the three-dimensional (3D) cell culture, the self-assembled peptides
PuraMatrixTM composed of standard amino acids (1% w/v) and 99% water were purchased
from Corning (New York, NY). The encapsulation of cells within the hydrogel scaffold
was prepared following the manufacturer’s instructions. Briefly, PuraMatrixTM stock
solution (1% w/v) was sonicated for 30 min in the ultrasonic bath to remove air bubbles
47
and reduce its viscosity. Then, the required amount of the PuraMatrixTM stock was
aliquoted and mixed with 20% sterile sucrose at a 1:1 ratio to yield a 0.5% w/v
concentration (half of the 20% sucrose was mixed with the gel and the other half with
suspended cells). At this point, 10 nm carboxy-functionalized superparamagnetic iron
oxide nanoparticles (Sigma-Aldrich, St. Louis, MO) were added at a concentration of 0.5%
of the total volume of the mixed solution. Subsequently, the cell suspension was prepared
by trypsinizing the flasks with a 0.25% trypsin/EDTA solution purchased from ScienCell
(Carlsbad, CA). Cells (1.5×105 cells/well) were resuspended in sterile 20% sucrose. The
cells/sucrose mixture was mixed equally with hydrogel mixture, to be transferred quickly
to the center of 24-well plates with total volume of 150 µL/well and 750 µL media was
added to form the gel in each well. Since the PuraMatrixTM pH is 2-2.5, the media was
changed twice to equilibrate to the physiological pH of 3D cell culture.
3.3 Cells viability assay
Lactate dehydrogenase (LDH) assay (CytoTox 96R Assay kit) purchased from Promega
(Madison, WI) was used to evaluate the cytotoxicity and proliferation of cells cultured
within the hydrogel. LDH is a stable cytosolic enzyme that can be measured from the cells
lysate, which reacts with a tetrazolium salt (i.e., iodonitrotetrazolium violet) to form a deep
red formazan dye. The number of cells was assessed at specific intervals (7, 14, and 21 d)
according to the manufacturer's protocol. Briefly, to prepare the lysis samples, the cells-
gel constructs were treated with a collagenase to digest the collagen secreted from cells.
48
The samples were incubated in 50 µL of collagenase for 30 minutes after washing with
PBS. Later, each sample was mixed with 500 µL of lysis buffer (1% Triton X-100), then
sonicated for 1 hour in a sonicator. The samples were then centrifuged at high speed. Next,
50 µL of the supernatant and 50 µL of CytoTox 96 Reagent were added to each well of a
96-well plate and covered with foil for 30 min at room temperature. Then 50 µL of the stop
solution was added to each well. Absorbance at 492nm was recorded and using the
calibration curve, cell number was quantified.
3.4 Cells differentiation assay
The alkaline phosphatase (ALP) assay was used to detect osteogenic differentiation of
hASCs. ALP is early osteogenic marker expressed on the cell surface. The SensoLyte®
pNPP Alkaline Phosphatase Assay Kit (AnaSpec, Fremont, CA) was used by measuring
the absorbance at 405 nm. Briefly, after removing the growth medium from each sample
and washing with PBS, 50 µL of collagenase was added and incubated for 30 minutes.
Then, 500 µL lysis buffer (1% Triton X-100) was added to each well. The cell lysate was
ready after sonication for 1 hour and centrifuging at 10,000 rpm for 5 min. Then 50 µL of
lysate cells were added to 96-well plate with a flat bottom, and 50 µL of pNPP substrate
solution was added to detect the ALP. Upon dephosphorylation of pNPP, the lysate samples
turn yellow. After incubation for 30 minutes at 37ºC, 50 µL of stop solution was added to
stop the reaction. With the absorbance reading from the microplate reader, the ALP
secretion was quantified using the calibration curve. Then, the data were normalized with
49
the total number of cells per well, expressing the ALP concentration in ng/cell. This assay
was performed at 7, 14, 21 d of incubation in triplicate.
3.5 Measurement of mineralization
Mineralization on the constructs was quantified using with Inductively Coupling Plasma-
OE Spectroscopy (ICP-OES, Varian) detecting calcium at 7, 14, and 21 d of incubation
(n=3). The samples were decalcified in 35% HCl (trace metal™ grade, Fisher Chemical),
followed by boiling for 5 h at 90ºC. Then, the samples were collected in liquid form, and
calcium was measured with ICP-OES through emissions at 396 nm. From the emission
intensity, the calcium concentration is obtained in mg/L (ppm) using a calibration curve.
3.6 Cells morphology
To investigate cell morphology in response to variations in the extracellular environment,
cytoskeletal actin microfilaments (F-actin) were stained using fluorescent phalloidin
conjugates (F-actin visualization Biochem kit). Also, 4',6-diamidino-2-phenylindole,
dihydrochloride (DAPI) was used to stain the cell’s nucleus This experiment was
performed at 3, 14 d and the results were visualized using by confocal microscopy
(Olympus FV1000) with an excitation/emission filter for F-actin (535/585 nm), and DAPI
(358/461 nm). The staining was done according to manufacturer's instructions. Briefly,
after removing media from each sample and washing them with washing buffer, 200µL of
50
the fixative solution was added for 10 min. Then, the samples were washed twice with
washing buffer to remove the fixative materials, 200 µl of permeabilization buffer was
added to each fixative sample. After 5 min, each sample was washed twice with washing
buffer. A 0.165 μM F-stain phalloidin and a 300 nM DAPI solution were prepared, then
the stained sample was covered with foil to keep it dark for 30 min at room
temperature. Then, each sample was washed three times to stop the reaction and covered
with mounting medium and coverslip.
3.7 Alkaline phosphatase staining
For qualitative investigation of cells differentiation, Stemgent Alkaline Phosphatase (ALP)
Staining Kit II (Cambridge, MA) has been used to detect ALP activity within the samples
at day 14. Briefly, after removing media, samples were washed with 0.05% concentration
of PBS containing Tween-20 as a permeabilizing agent. 0.5ml of fixative solution have
been added for 5 to 10 min. Then, the samples were washed twice, and 0.6 ml of ALP
staining was added that composed of equal ratio 1:1:1 of AP substrate solution (a mixture
of 0.2 ml of solution A and 0.2 ml of solution and 0.2 ml of solution C) for 15 minutes at
room temperature. For stopping the reaction, the samples washed twice with PBS. Later,
the samples were covered with mounting medium and coverslip to prevent drying. By
using, the ALP expression was detected as a red or purple stain.
51
3.8 Fourier transform infrared spectroscopy (FTIR) analysis
The self-assembled peptides (PuraMatrix) with and without iron oxides nanoparticles
scaffolds have been characterized by using FTIR device. Since FTIR spectrum provides
information about specific functional groups presented in each of hydrogel type through
measuring the transmittance frequency at which specific atoms will be vibrating. The
characteristic of testing hydrogel will show the transmittance and wavelength that range
between 400 to 4000 cm– 1.
3.9 Electromagnetic fields exposure system
The equipment that used to generate extremely low-frequency pEMFs consisted of a
function generator, oscilloscope, Helmholtz coils, and µ-metal box. The function generator
(Agilent) generated the burst signal with specific frequency, amplitude, and shape. The test
signal was equivalent to FDA approved signal for bone fracture healing, which consists of
15 Hz pulse burst of 20 pulses with magnetic field increased from 0 to 1 mT in 5 ms and
then decreased to zero in 61 ms [Bassett, et al., 1982; Daish, et al., 2018, Polo-Corrales, et
al., 2018]. An oscilloscope was used to display the generated signals in terms of voltage as
a function of time. The EMF was generated using Helmholtz coils (3B Scientific®
Physics), which consist of a pair of copper coils that operate on alternating fields. When an
alternating current ran through the Helmholtz coils, uniform electromagnetic fields are
generated in space over a considerable volume. The samples were placed in the center of
52
the coils. The main feature of each coil is composed of 124 turns, outer coil diameter is
311 mm, inner coil diameter is 287 mm, mean coil radius is 150 mm, and coil resistance is
1.2 ohm. Thus, from coil featured and the current intensity, the magnitude of magnetic flux
density (B) can be measured according to the following formula (Gupta, et al., 1991):
𝐵 = (5
4)
32
𝜇0𝐼𝑛
𝑅
where n is the number of turns in each coil, R is mean coil radius and 𝜇0 is permeability
of free space (4π × 10−7 Tꞏm/A). which leads to B = 7.433×10‾⁴ I in T.
The Helmholtz coils were placed in the incubator inside a μ–metal enclosure, which shields
against the earth static fields and low magnetic fields from the equipment around it. This
assures that the only EMFs exposure comes from the Helmholtz coils. This μ–metal is
composed mainly of nickel, iron, and some copper or chromium that gives a low reluctance
path for the magnetic flux.
The experimental groups were divided into two, pEMF stimulated and non-stimulated.
Each experimental group consisted of four different formulations, all based on hASC-
seeded scaffolds, with/without osteogenic induction media, and with/without NPs. The
stimulated group was exposed to pEMFs for 8 h every day. The non-stimulated group was
placed in the pEMF system for 8 h per day but it was kept off. This accounted for any
differences in cell culture conditions between the pEMF system, and the cell culture
incubator, where all cell cultures were kept the remaining time.
53
Figure 6. Electromagnetic exposure system: a) closed mu-metal box and b) Helmholtz
coils with cells.
3.10 Statistical analysis
All numerical data were evaluated statistically according to the ANOVA test followed
Dunnett test to identify the significant differences. Values of p < 0.05 were accepted to be
used as a significant level of difference between the mean of the experimental and
corresponding control groups.
a b
54
CHAPTER IV
Results &Discussion
4.1 Results & Discussion
4.1.1 Viability and Proliferation
Cell proliferation was quantified using an LDH cytotoxicity assay using passage 4
hASCs for four different formulations that were either pEMF stimulated or non-stimulated.
The results were assessed after 7, 14, and 21 days of culture for all formulations. The results
showed (Figure7) that cells start to grow rapidly for all groups. However, there was no
statistically significant difference in cellular viability between groups with and without
pEMF stimulation.
At day 7, within the non-stimulated (i.e., without pEMF or osteogenic induction
media) group there was a significant increase (p<0.001) in the number of viable cells
detected within the superparamagnetic (NP-doped) scaffold over those using the hydrogel
alone. This revealed that the carboxy-functionalized iron oxide NPs not only do they have
any cytotoxic effect on the cells but in addition promote their proliferation. It is known that
carboxy functionalities increases hydrophilicity and significantly increases the
proliferation and osteogenic differentiation of MSCs (Phillips, et al., 2010). It has also been
shown that by using electrospun paramagnetic scaffolds based on γ-Fe2O3/nano-
hydroxyapatite/poly (lactic acid), the proliferation of preosteoblasts was enhanced
55
the group of superparamagnetic scaffolds under pEMF stimulation (without osteogenic
induction media) showed a similar proliferation rate to the non-stimulated group, which
were significantly different (p<0.001) to the control (lacking NPs). The biochemical
stimulation using osteogenic induction media alone was enough to significantly (p<0.001)
enhance cell proliferation when compared to the non-stimulated control. Although this
significant increase in the rate of proliferation was also seen on both pEMF stimulated and
non-stimulated superparamagnetic scaffold groups, there was no significant difference
between them, but only with respect to the control (p<0.0001).
At day 14, all the groups continue to grow, under basal media, with similar
proliferation rates, but with no significant differences between pEMF stimulated and non-
stimulated groups regardless of the presence of NP-doped superparamagnetic scaffolds,
when compared to the control. When cells were stimulated with osteogenic induction
media there were no significant differences between pEMF stimulated and non-stimulated
groups when grown in the hydrogel alone, but significantly different with the control
(p<0.05). On the other hand, there was a significant difference (p<0.001) when these were
grown in superparamagnetic scaffolds, with enhanced the cellular proliferation with pEMF
stimulation. At day 21, we observe the same trend as for day 14. In addition, the group of
stimulated with osteogenic induction media has a significant difference (p<0.05) as
compared to control.
56
Figure 7. LDH assay. The proliferation of hASCs seeded within self-assembled peptide
hydrogel with and without NPs under extremely low-frequency pEMF of 1 mT was
quantified, after 7, 14, 21 d. *p < 0.05; **p < 0.001; ***p < 0.0001, indicate statistically
significant differences between basal media group (control) and the other groups. The
mean values are calculated from the average results of three samples, the results are
represented as mean ± SD.
In this study, the results from the proliferation assay showed no cytotoxicity from
the hydrogel with or without the presence of NPs towards hASCs, showing an appreciable
proliferation rate. The assessment of biochemical, mechanical, and electromagnetic
stimulation on hASC proliferation shows that the simultaneous presence of osteogenic
induction media and superparamagnetic NPs has the strongest effect on cell proliferation,
without any significant impact by pEMFs. This lack of effect due to pEMF stimulation on
proliferation was also seen in the non-stimulated (i.e., biochemical and mechanical)
57
controls. The presence of either osteogenic induction media or superparamagnetic NPs
had an equivalent positive effect on hASC proliferation. Although this positive effect was
significantly less than the combined presence of osteogenic induction media and NPs. Also,
during the time and after 21 days of culture the cells have the same trend of proliferation
as the cells grown after 14 days. This behavior is due to cells that differentiated and stopped
proliferating, whereas cells exhibit cell-cycle arrest following differentiation by the
activation and deactivation of a collection of cyclin-dependent kinases, which regulate
specific steps in the cell cycle (Myster, et al., 2000).
The objective of adding superparamagnetic NPs to the hydrogel was so that the
matrix would be induced to mechanically vibrate with the application of an alternating
magnetic field (Golovin, et al., 2017). The results show that although the presence of
superparamagnetic NPs promotes hASC proliferation, the mechanical vibration induced by
the pEMFs is either not required or that the force generated is too weak to elicit a response
from the cells. It must thus be an intrinsic property of the NPs, which stimulates cell
proliferation and that further boosts this effect in the presence of osteogenic induction
media. This is consistent with published results showing the osteoinductive effects of
superparamagnetic NPs without the application of external magnetic fields (Wu, et al.,
2010; Wei, et al., 2011; Yun, et al., 2015). Although many of the studies reporting this
effect do not provide possible mechanisms to explain this phenomenon, some have
explored different hypotheses (Castro, et al., 2017; Zhu, et al., 2017). One of these
hypotheses demonstrates that the influence of the superparamagnetic NPs within the
scaffold at the nanoscale level act as a single magnetic domain, initiating micromotions.
58
This movement is translated by the cells as a trigger for the mechanotransduction pathway
by affecting the ions channels (Gil, et al., 2014). Other studies demonstrate the effect of
superparamagnetic NPs on cell proliferation by significantly decreasing the intracellular
H₂O₂ and peroxidase-like activity. They also found that NPs had an impact on cells cycle
through lysosomal metabolism of iron oxide particles that led to iron depletion, which
might lead to cell cycle block at G1/S, and it affects the expression of regulating molecules
vital for the cell cycle process and apoptosis (Huang, et al., 2009). Furthermore, a very
thorough and interesting study proposed that the composition of the protein corona that
forms on the superparamagnetic NPs could provide the necessary stimulation to promote
the levels of proliferation measured. The study characterized the protein composition of
the corona after being exposed to fetal bovine serum alone or in the presence of the proteins
secreted by pre-osteoblasts (Zhu, et al., 2017). Their findings showed the presence of
proteins related to calcium ions, G-protein coupled receptors, and MAPK/ERK cascades
as compared with scaffolds not containing NPs. All these could be related to helping in the
induction of cell proliferation and could thus explain the increase in hASC proliferation in
the presence of superparamagnetic scaffolds without pEMF stimulation. This hypothesis
supports the proliferation results (Figure 7) since the groups that were cultured in the
presence of osteogenic media and superparamagnetic scaffolds showed a high proliferative
rate due to the formation of protein corona composed by the components of the osteogenic
media allowing these to have a higher stability and longer half-life.
It has been shown that the force necessary to stimulate of a single ion channel is of
the order of 1-2 pN (Yoshimura, et al., 2010). Other studies have demonstrated that a
mechanical force less than 0.2 pN is required to activate the TREK-1 channel and of 2 pN
59
to break the bond between fibronectin and the cytoskeleton (Sapir-Lekhovitser, et al.,
2016). However, our calculations showed that the magnetic force on our systems was about
0.148 pN per cell and its influence was noticed for a group of magnetic scaffolds under
pEMF during the first week of culture to have a significant degree (p<0.001) as compared
to the control. Nevertheless, after 14 and 21 d the magnetic force was not strong enough to
activate the mechanotransduction process due to the larger number of cells as shown by
the group of magnetic scaffolds under pEMF that did not exhibit significant changes as
compared to the control. [APPENDIX]
4.1.2 Differentiation to osteoblasts
Osteoblast differentiation from hASCs was detected measuring the activity of the
ALP biomarker after 7, 14, 21 d of culture (Wang, et al., 2007). As before, the two
experimental groups were tested, pEMF stimulated and non-stimulated, which were
subdivided into four different formulations, all based on hASC-seeded scaffolds,
with/without osteogenic induction media, and with/without NPs.
ALP activity at day 7 showed signs of early differentiation of hASCs cells into
osteoblasts (Figure 8). However, the cells cultured within the group of superparamagnetic
scaffolds under pEMF stimulation (without osteogenic induction media) showed a
statistically similar low level of ALP activity as the non-stimulated group and the controls
(i.e., with and without pEMF stimulation and both lacking NPs). However, when the
hASCs in hydrogels (no NPs) were simultaneously stimulated with osteogenic induction
media and pEMF, demonstrated signs of an early differentiation to osteoblasts with a
60
significant increase in ALP activity (p<0.05) when compared to the control. In contrast, a
lack of pEMF stimulation showed low levels of ALP activity with no significant change
when compared to the control. Under these conditions, the presence of superparamagnetic
scaffolds shows high levels of ALP activity with significant difference (p<0.001) without
pEMF stimulation and even higher with pEMF stimulation significantly higher than the
control (p<0.0001).
Although after 14 d of culture, the levels of ALP activity had significantly increased
in all groups, those using superparamagnetic scaffolds (without osteogenic induction
media) with and without pEMF stimulation showed no significant differences when
compared to the control. In the case of the group under pEMF stimulation (without NPs or
osteogenic induction media) had a slight increase of ALP activity, it was not significant
when compared to the control.
However, when cells were stimulated with osteogenic induction media (not pEMF)
there was a significant increase in differentiation as measured by ALP activity (p<0.001),
with the addition of pEMF stimulation showing a significant increase. As with the results
at day 7, the presence of superparamagnetic scaffolds led to elevated levels of ALP activity
with significant difference (p<0.001) without pEMF stimulation and even significantly
higher with pEMF stimulation (p<0.0001).
At day 21, we still observe an increase in the ALP activity with time, and we noticed
that the groups with osteogenic induction stimulation regardless of the presence of pEMF
stimulation were significantly different from the control (p<0.0001). Similarly, the
presence of superparamagnetic scaffolds, regardless of the presence of pEMF stimulation,
were significantly different from the control (p<0.0001). Moreover, ALP activity has
61
significantly increased (p<0.05) for the first time for the superparamagnetic scaffolds group
without osteogenic induction media stimulation, but with pEMF stimulation, while the
other groups remain unchanged.
The results from ALP assay showed that the hASCs start to differentiate into the
osteoblast phenotype since day 7 when grown in the superparamagnetic scaffold and
cultured with osteogenic induction media. Also, early differentiation at day 7 was noticed
in the group cultured within osteogenic media, with and without pEMF stimulation. With
the ALP activity being higher with the latter. These effects were not seen in the presence
of osteogenic induction media alone, but only after 14 d of stimulation. At this same time,
it can also be seen that hASCs biochemically stimulated with induction media continue to
increase their level of differentiation, with a non-significant increase due to the presence
of pEMF stimulation and a non-significant difference due to the presence of NPs. After 2
weeks a trend starts to emerge which involves a significant level of osteogenic
differentiation by hASCs stimulated with osteogenic induction media, with a non-
significant increase due to pEMF stimulation and no effect due to the superparamagnetic
scaffolds. We also noticed that ALP activity, an early osteogenic marker, did not show a
significant difference between groups with/without magnetic fields after 21 days of culture.
This could be since ALP is not a good late osteogenic biomarker. Perhaps the use of a late
osteogenic biomarker such as osteocalcin is a better choice (Granéli, et al., 2014), to show
significant differences between groups.
If the proposed NP protein corona mechanism (Zhu, et al., 2017) is correct, then
the significant difference in ALP activity see after 7 days in the presence of
superparamagnetic scaffolds is due to osteogenic proteins adsorbed onto the NPs and
62
inducing hASC differentiation. During this initial stage, there is a sustained presence of
osteogenic proteins which allows an early commitment of hASC to the osteogenic lineage.
After lineage commitment, there is less need for osteogenic proteins to continue the
differentiation process (Ferroni, et al., 2018). So, the presence of the superparamagnetic
scaffolds helps improve the levels of hASC osteogenic differentiation by allowing an early
commitment of these cells to the osteogenic lineage.
Figure 8. Alkaline phosphatase activity. The differentiation of hASCs cells was assessed
at 7, 14, 21 d. ALP values were normalized with the number of cells of each sample. *p <
0.05; **p < 0.001; ***p < 0.0001, indicate statistically significant differences between
basal media group (control) and the other groups (one-way ANOVA followed by Dunnett
test). The mean values are calculated from the average results of three samples, the results
are represented as mean ± SD.
63
4.1.3 Mineralization
While ALP is relatively an early differentiation marker that increases during the
proliferation and matrix synthesis stage (Harris, 1990), the matrix calcium deposition
defines the terminal stage in osteoblast maturation (Cormier, 1995). Therefore, the series
of osteogenic differentiation assays, involving ALP activity and staining, and the calcium
mineralization, clearly demonstrate the significant role of the 3D superparamagnetic
scaffolds played in accelerating osteoblastogenesis of hASCs. The calcium depositions
were quantified for all the groups over three weeks (at day 7, 14, and 21) as shown in
Figure 9.
At day 7, results revealed that four groups did not exhibit any early calcium
depositions, these included the hASC-seeded scaffolds, with pEMF stimulation, with
osteogenic induction media stimulation, and NP-containing scaffolds without any
stimulation. However, the group with NP-containing (superparamagnetic) scaffolds were
stimulated with induction media showed some early calcium deposition. Also, although the
remaining three groups, which were all pEMF-stimulated, showed some early calcium
deposition, it was not significantly different from the superparamagnetic scaffold only
stimulated with osteogenic induction media. These groups were superparamagnetic
scaffolds with and without induction media and hydrogel (no NPs) with induction media,
all three being pEMF-stimulated.
After 7 d, non-significant traces of calcium in the extracellular matrix is observed.
This is true for groups with pEMF stimulation or osteogenic media and no pEMF. After 14
64
d, even the hASC-seeded control (without stimulation) start to show traces of calcium in
the extracellular matrix. Meanwhile, the groups of superparamagnetic scaffolds stimulated
with induction media and with and without pEMF stimulation, showed higher levels of
calcium deposition, but not significantly higher than the control. After 21 d, there was a
significant increase (p<0.05) in calcium deposition, in the two groups with
superparamagnetic scaffolds being pEMF-stimulated, with and without osteogenic
induction media stimulation. The remaining groups deposited equivalent amounts of
calcium to the control.
Groups with a low initial (day 7 and 14) ALP activity did not form lots of
mineralized matrix later in day 21, especially the control group. However, mineralization
measurements support the ALP activity results, whereas the samples that were cultured in
superparamagnetic scaffolds and osteogenic induction media showed an early calcium
deposition after 7 d. In fact, this effect can be present during the first 2 weeks. These are
the groups with early ALP increase.
In addition, there is a clear and significant increase in mineralization after 21 d due
to pEMF stimulation. As the results show, the hASCs have already differentiated and do
not need further stimulation from the induction media or NPs and are depositing mineral
stimulated solely by pEMFs. A group with osteogenic media and nanoparticles had high
ALP from the very initial culture (day 7) and showed non-significant mineralization on day
7 compared to the control. We believe that the cells in this group are differentiated very
early and secreting calcium earlier than groups with pEMF stimulation, so the ECM
calcium is comparable to the group with pEMF stimulation on day 21.
65
.
Figure 9. Mineralization assay. Calcium concentrations were quantified after 7,14, 21
days. Error bar represents the SD. p < 0.05, indicate statistically significant differences
between MSCs media gel+cells group (control) and the other groups (one-way ANOVA
followed by Dunnett test). The mean values are calculated from the average results of three
samples, the results are represented as mean ± SD.
4.1.4 Cell morphology
A qualitative test used to test cell morphology is through of actin and DAPI
staining, as shown in Figure 10, at day 3 and 14. At day 3, after culturing the cells within
the hydrogel, it got a spherical shape at an early stage especially when they were cultured
with mesenchymal media either with and without exposing to magnetic fields (Figure 10;
-2E-08
0
2E-08
4E-08
6E-08
8E-08
0.0000001
1.2E-07
1.4E-07
1.6E-07
7d 14d 21d
No
rmal
ized
Cal
ciu
m c
on
cen
trat
ion
mg/
L (p
pm
)
Time (day)
MSC media (gel+cells)
Msc media (gel Np+ cells)
osteo media (gel+cells)
osteo media (gel Np+cells)
MSC media (gel+cells)+MF
Msc media (gel Np+ cells)+MF
osteo media (gel+cells)+MF
osteo media (gelNp+cells)+MF
P<0.05
66
Day 3; A, B, E, F). Also, at the same time point when we used an induction media, the cells
have a spindle and elongated shape morphology, and this was noticed with or without
applying magnetic fields which is the sign of early osteoblastic differentiation (Figure 10;
Day 3; C, D, G, H). Others have shown the same where the elongated cells were the sign
of differentiation or specialization into osteoblasts after following adherence to the scaffold
(Kannarkat et al., 2010). At Day 14, the images showed that the cells were entirely
embedded in the nanofiber hydrogel by taking the shape of the fibers (Figure 10; Day 14).
67
Figure 10. Phalloidin-labeled actin filaments stain (red) and DAPI stain (blue) for hASCs
within hydrogel at 3 and 14 days. A) basal media, B) basal media + NPs, C) osteogenic
media, D) osteogenic media + NPs, E) basal media + pEMF, F) basal media + NPs + pEMF,
G) osteogenic media + pEMF, H) osteogenic media + NPs + pEMF. Scale bars at day 3
and 14 are 100 µm (10X) and 50 µm (20X), respectively. Biochemical stimulation –
osteogenic induction media.
The results on hASC morphology showed that the cells intimately interacted with
the ECM since the scaffold is mainly composed of water, allowing the cells to freely
68
migrate and interact. Also, the presence of NPs or EMFs did not affect cell morphology.
During the early days of seeding hASCs adopted a spherical shape, but later they started to
elongate and got spindle morphology. Fan, et al., mention that the hASCs need to adhere
to a surface to allow them to stretch and to get its original shape to be able to contact with
other cells, if not they will become apoptotic (Fan and Wang, 2017).
4.1.5 Alkaline phosphatase staining
To do a qualitative evaluation of hASCs osteoblast differentiation, ALP staining
was used at day 14 for all eight groups as shown in Figure 11. The results demonstrate that
at day 14, hASCs had been differentiated into osteoblasts. Whereas the groups that were
cultured with and without NPs, and without pEMF or induction media stimulation showed
only marginal ALP staining (Figure 11, A and B), when these were pEMF-stimulated, they
showed high levels of ALP staining (Figure 11, E and F). In addition, the results showed a
higher level of ALP staining when stimulated with osteogenic induction media, with or
without pEMF stimulation (Figure 11, C, D, G, and H).
69
Figure 11. Alkaline phosphatase stain for hASCs cells seeded within hydrogel at day 14,
scale bar = 200 µm (10X). A) basal media, B) basal media + NPs, C) osteogenic media, D)
osteogenic media + NPs, E) basal media + pEMF, F) basal media + NPs + pEMF, G)
osteogenic media + pEMF, H) osteogenic media + NPs + pEMF. Biochemical stimulation
– osteogenic induction media.
As expected, the ALP staining result was consistent with the quantitative ALP
activity results. At day 14, biochemically-stimulated cells contained high levels of alkaline
phosphatase, with a non-significant increase due to the presence of pEMF stimulation and
a non-significant difference due to the presence of NPs.
70
4.1.6 FTIR analysis
FTIR analysis of the self-assembled peptides scaffolds with or without iron oxide
nanoparticles showed that the presence of nanoparticles did not cause much difference in
the spectrum from the scaffold without nanoparticles. As shown in Figure 12, the hydrogel
with nanoparticles has shown three major peaks, located at (3625-3100), 2100, 1625 cm–1.
The spectrum showed that there is a broad –OH stretching band between 3625 and 3100
cm–1, which is mainly resulted from water crystal and characteristic of the O-H stretch band
of the hydroxyl group. A medium peak is shown on 2100 cm–1 that represent the C≡C bond.
Also, the spectrum displays at 1625 cm–1 peak which is the characteristic of amide I bond
that link the amino acids of the hydrogel which resulted from stretching vibrations of the
C=O bond of the amide, which represented the existence of the β-sheet structures of the
hydrogel. Also, FTIR spectrum for hydrogel alone showed almost the same spectrum with
a slight shift. However, there is a strong stretching band of –OH at 3400 cm–1, and two
medium peaks 2400 and 2100 cm–1 to represent the C≡C bond. Also, another peak in 1620
to represent the amid I of stretching vibrations of the C=O bond.
71
Figure 12. FTIR analysis of self-assembled peptides with and without iron oxide
nanoparticles.
4.2 Conclusion
In this study, we developed a superparamagnetic scaffold based on a biomimetic
hydrogel for the 3D culture of hASCs to study their response to biochemical,
electromagnetic and mechanical stimulation. This was done by evaluating proliferation,
osteoblastic differentiation, ECM mineralization, and morphology under a combination of
these forms of stimulation. The 3D superparamagnetic scaffold was based on peptides that
self-assembled to form nanofibers and an ECM-type structure that allows nutrients and
oxygen to be effectively transported to the seeded cells in a manner to the natural condition
(Zhang, 2004). The specific peptide used was 16-mer that consisted of four repeats of the
RADA amino acid sequence. Hydrogels made from this peptide have demonstrated the
72
capacity to enhance the proliferation and differentiation of primary osteoblasts in vitro
(Bokhari, et al., 2005) and in vivo (Misawa, et al., 2006). The choice of hASC over bone
marrow-derived MSCs is based on their capacity to proliferate faster and retain and
enhanced longer an enhanced capacity for differentiation over their bone marrow
counterparts (Burrow, et al., 2017)
Our results show no negative or cytotoxic effects on hASCs, due to the presence of
pEMFs or superparamagnetic NPs. Also, the result showed that there was an early
differentiation by incorporating osteogenic media either with the presence of
superparamagnetic NPs or pEMFs. The principal results of the present study revealed
several novel findings regarding the events involved in the induction of the osteogenic
differentiation of hASCs by pEMFs and mechanical stimulation. For instance, that adding
superparamagnetic NPs to the hydrogel induced a significant increase in proliferation, but
not due to mechanical stimulation due to NP vibrations induced by the pEMFs. We propose
that this effect is due to the proteins adsorbed onto the NPs, which help induce cell
proliferation. Although a positive effect due to mechanical stimulation was expected,
perhaps the use of extremely low-frequency field did actuate a level of vibration in the
superparamagnetic NPs to mechanically stimulate the hASCs (Golovin, et al., 2017). In
addition, it was seen that after two weeks there was a significant level of osteogenic
differentiation by hASCs stimulated with osteogenic induction media, with a non-
significant increase due to pEMF stimulation and no effect due to the superparamagnetic
scaffolds. Perhaps, prior to that due to a sustained presence of osteogenic proteins on the
corona of the NPs, there was an early commitment of hASC to the osteogenic lineage.
Committed cells then became significantly more responsive to pEMF stimulation not only
73
promoting osteogenic differentiation, as evidenced by ALP activity and staining, the extent
of ECM mineralization (Ferroni, et al., 2018).
74
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89
Appendix I
To estimate the magnetic force generated on each cell to activate the mechanotransduction
process we decided to calculate the total force and divide it by a number of cells in each
well plate. The magnetic force (F) acting on magnetic NPs inside a magnetic field which
is defined as:
𝐹 = 𝑉∆𝜒
µₒ(𝐵. 𝛻)𝐵
where 𝑉 is the NP volume (in m3), ∆χ is the difference in magnetic susceptibilities between
the NPs and the surrounding medium (dimensionless), µₒ is the permeability of vacuum,
which is a constant equal to 4π×10-7 T·m/A, B is the applied magnetic field (in T), (𝐵. 𝛻)
is the gradient of the magnetic field (in T/m).
The magnitude of magnetic flux density (B) generated from the Helmholtz coils was
measured from coil featured and the current intensity, can be measured according to the
following formula:
𝐵 = (5
4)
32
𝜇0 𝐼 𝑁
𝑅
where N is the number of turns in each coil (N = 124), R is mean coil radius (R = 150 mm),
I is the electric current passing through the Helmholtz coil to generate magnetic fields (I =
1.345 A); thus, the magnetic field intensity is
90
B = (5
4)
3
24𝜋 × 10−7 × 1.345 ×
124
0.15 = 1 mT = 1 × 10−3 T
The expression for the gradient of the magnetic field at the center of coils carrying
current in the reverse direction, this gradient is given by:
(𝐵. ∇) =ⅆ𝐵
ⅆ𝑥|
𝑥=0=
3
2(
4
5)
52
⋅𝑁𝜇0𝐼
𝑅2
where N is the number of turns in each coil (N = 124), R is mean coil radius (R = 150 mm),
I is the current (I = 1.345 A).
(𝐵∇) =ⅆ𝐵
ⅆ𝑥|
𝑥=0=
3
2(
4
5)
5
2 124×4𝜋×10−7× 1.345
(0.15)2 = 0.007998 T/m
To measure superparamagnetic iron oxide NP volume (V), depending on the manufacturer
information, each nanoparticle has an average diameter of d = 10 nm. So,
V =4
3𝜋𝑟3 =
4
3𝜋(5 × 10−9)3 m3/NP
To quantify the difference in magnetic susceptibilities ∆χ we need to have the NPs and the
surrounding medium (dimensionless) susceptibilities. The susceptibility of NPs (is
dependent on the frequency of the magnetic field) χNP = 0.115 (Grüttner, et al., 2007), and
susceptibility of media χ0 = 0.
Thus, according to the previously calculated information, the magnetic force generated
from a single NP is given by
91
𝐹 = 𝑉∆𝑥
µₒ(𝐵𝛻)𝐵 =
4𝜋
3(5×10−9
)3
×0.115×0.007998×10−3
4𝜋×10−7
F = 37.36× 10−25 N/NP = 37.36 × 10−13
pN/NP
To estimate the number of NP per well, we will use the manufacturer information about
NP density (1 g/ml), NP concentration (5 mg/mL). The NPs concertation added during
magnetic scaffold synthesis was 60 µg/ml, so the NP mass per well is calculated as shown
below:
Mass of NPs in each well = 60×0.5
100×103 = 3 × 10−3 g/well
The total volume of NPs in each well = mass
ⅆ𝑒𝑛𝑠𝑖𝑡𝑦=
3×10−3 𝑔
1 𝑔/𝑚𝑙= 3 × 10−3
ml =
3 × 10−9 𝑚ᶟ/well
Thus, the
Total number of NPs per well = 3×10−9
𝑚ᶟ4
3𝜋(5×10−9 )3 𝑚ᶟ
= 6 × 1015 NPs/well
Using the number of cells seeded at day 0 as 15×104 cells, and from the total number of
NPs per well, we can estimate the total magnetic force generated from the NPs around each
cell considering that NPs are not affecting other cells:
Number of NPs around each cell = total number of NPs per well
𝑡𝑜𝑡𝑎𝑙 𝑛𝑢𝑚𝑏𝑒𝑟 𝑜𝑓 𝑐𝑒𝑙𝑙𝑠 𝑝𝑒𝑟 𝑤𝑒𝑙𝑙=
6×1015
15×104 =
0.4 × 1011 NPs/cell
Therefore, the total force generated from the NPs around each cell will be
F = 0.4 × 1011 × 37.36 × 10−13 pN/cell = 0.148 pN/cell
This is a conservative calculation simply because we cannot calculate the exact resultant
force vector on each cell.