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Bone Regeneration: From Intramembranous
to Endochondral Pathway
Wanxun Yang
Colofon
Thesis Radboud University Medical Center, with summary in English and Dutch.
Bone Regeneration: From Intramembranous to Endochondral Pathway
The research presented in this thesis was financially supported by China Scholarship Council
(No. 2010627030) and The Royal Netherlands Academy of Arts and Sciences (KNAW).
Publication of this thesis was financially supported by Radboud University Medical Center
(Radboud UMC), Radboud Institute for Molecular Life Sciences (RIMLS), Netherlands
Society for Biomaterials and Tissue Engineering (NBTE) and Corbion Purac Biomaterials.
Layout:
Arnold Nijhuis & Wanxun Yang
Cover design:
Wanxun Yang & Proefschriftmaken.nl || Uitgeverij BOXPress
Printed by:
Proefschriftmaken.nl || Uitgeverij BOXPress
Published by:
Uitgeverij BOXPress, ‘s‐Hertogenbosch
ISBN: 978-90-8891-962-6
Copyright © Wanxun Yang, 2014
All rights reserved.
Bone Regeneration: From Intramembranous to
Endochondral Pathway
Proefschrift
ter verkrijging van de graad van doctor
aan de Radboud Universiteit Nijmegen
op gezag van de rector magnificus prof. dr. Th.L.M. Engelen,
volgens besluit van het college van decanen
in het openbaar te verdedigen op dinsdag 21 oktober 2014
om 14.30 uur precies
door
Wanxun Yang
Geboren op 15 juni 1985
te Bengbu, China
Promotor
Prof. dr. J.A. Jansen
Copromotoren
Dr. F. Yang
Dr. ing. S.K. Both (Universiteit Twente)
Manuscriptcommissie
Prof. dr. P. Buma
Prof. dr. J. Klein-Nulend (Vrije Universiteit)
Prof. dr. M. Karperien (Universiteit Twente)
Paranimfen
Eva Urquía Edreira
Alexey Klymov
Bone Regeneration: From Intramembranous to
Endochondral Pathway
Doctoral Thesis
to obtain the degree of doctor
from Radboud University Nijmegen
on the authority of the Rector Magnificus prof. dr. Th.L.M. Engelen,
according to the decision of the Council of Deans
to be defended in public on Tuesday, October 21, 2014
at 14.30 hours
by
Wanxun Yang
born on June 15, 1985
in Bengbu, China
Supervisor
Prof. dr. J.A. Jansen
Co‐supervisors
Dr. F. Yang
Dr. ing. S.K. Both (University of Twente)
Doctoral Thesis Committee
Prof. dr. P. Buma
Prof. dr. J. Klein-Nulend (VU University)
Prof. dr. M. Karperien (University of Twente)
Ushers
Eva Urquía Edreira
Alexey Klymov
“Don’t be too timid and squeamish about your actions.
All life is an experiment.”
- Ralph Waldo Emerson
To my beloved family
谨以此书献给我的父母和家人
Contents
Chapter 1 | Page 12
General introduction
Chapter 2 | Page 20
Combining osteochondral stem cells and biodegradable materials for bone
regeneration
Chapter 3 | Page 40
Biological evaluation of porous aliphatic polyurethane/hydroxyapatite composite
scaffolds for bone tissue engineering
Chapter 4 | Page 60
In vivo bone generation via the endochondral pathway on three-dimensional
electrospun fibers
Chapter 5 | Page 80
Performance of different three-dimensional scaffolds for in vivo endochondral
bone generation
Chapter 6 | Page 106
Effects of in vitro chondrogenic priming time of bone marrow derived mesenchymal
stem cells on in vivo endochondral bone formation
Chapter 7 | Page 130
Roles of cartilage templates in mediating mesenchymal stem cell behavior in
endochondral bone tissue engineering
Chapter 8 | Page 148
Summary, closing remarks and future perspectives
Chapter 9 | Page 158
Samenvatting, slotopmerkingen en toekomstperspectieven
Acknowledgements | Page 168
List of publications | Page 176
Curriculum Vitae | Page 178
Chapter 1
General Introduction
General Introduction
1. Background
Bone is a dynamic and vascularized tissue responsible for providing structural
support to the body and protection of the internal organs. In addition, it plays a
key role in supporting muscle contraction, blood production and the storage of
mineral components. Clinically, the occurrence of bone defects is common and
can be caused by a variety of reasons, including congenital deformity, trauma/
injury, infection and tumorectomy. Although bone possesses a remarkable ability
to heal and remodel, in many clinical cases the regeneration of bone defects
exceeds its self-healing potential. It has been reported that bone is the second
most transplanted tissue after blood with over 2.2 million bone grafting procedure
being performed annually worldwide [1]. In addition, the treatment for fractures
and other musculoskeletal injuries will continue to boost as a result of increasing
life expectancy and dynamism in daily activities.
Current clinical treatments for bone defects involve the transplantation of autograft
(from the patient) or allograft (from the donor), which can be harvested from the
iliac crest, fibula, scapula or radius [2]. So far, this approach has been established
as the “gold standard” for bone reconstructive surgery. However, the inherent
drawbacks of this approach, such as limited source of available tissue, donor-site
morbidity, risk of rejection or disease transmission (when using allograft), and
relevant surgical complications, hinder its applications in the clinical practice,
especially for the treatment of large bone defects [3, 4]. This has strongly urged the
clinicians and researches to explore alternative treatment options.
2. Bone tissue engineering
Since the concept of tissue engineering has been introduced by Langer and
Vacanti [5], it has emerged as a promising approach for regenerating bone tissue.
The essential rationale behind this is to utilize the body’s own capacity for bone
healing in conjunction with engineering principles. Three important elements
are involved in bone tissue engineering, i.e. scaffolds, cells and biomolecules. A
conventional tissue engineering strategy starts with loading progenitor cells onto
proper designed biomaterials to induce cell differentiation in a specific pathway,
followed by transplantation of this cell-scaffold complex into the defect site to
regenerate bone tissue. In this procedure, cells are the main source to form the
new bone tissue, among which the progenitor/stem cells isolated from bone
marrow have been recognized as the most efficient cell source [6].
15
1
Chapter 1
3. Intramembranous and endochondral approach for bone
tissue engineering
During osteogenesis, bone is essentially formed through two different mechanisms:
intramembranous ossification and endochondral ossification. Accordingly, this
has inspired researchers to implement bone tissue engineering through two
specific pathways of cell differentiation. In the first approach, bone is directly from
osteoprogenitor cells in a manner akin to intramembranous ossification [7]. This
method is the most commonly used strategy in bone tissue engineering for years,
a standard procedure of which involves the differentiation of osteoprogenitor
cells into osteoblasts. The second approach resembles endochondral ossification,
whereby a cartilage intermediate is created and used as an osteoinductive
template to transform into bone in vivo [8]. Such endochondral approach has been
recently proposed and shows superiority to generate vascularized bone tissue, as
cartilage is hypoxia tolerant and chondrocytes can secrete angiogenic factors to
promote vascular invasion into the engineered constructs [9, 10].
4. Scaffolds for bone tissue engineering
Besides the cells, the choice of biomaterial as scaffold is also critical to the success
of tissue engineering. It provides a three-dimensional environment for the cells
and plays an imperative role in regulating the cellular behaviors, i.e. proliferation
and differentiation [11]. The shape and architecture of the scaffold define the
ultimate structure of newly formed tissue [12].
Since the introduction of the tissue engineering concept, a variety of biomaterials
of natural derivation or synthetic origin has been suggested as potential scaffold
candidates. Among them, ceramic-based materials, such as hydroxyapatite (HA)
and calcium phosphate (CaP), and polymer-based materials, such as collagen,
polycaprolactone (PCL), polyurethanes (PUs), polyglycolide (PGA), polylactide
(PLA) and their copolymers, are currently the most widely used. Along with
research on biomaterials, techniques to manufacture scaffolds have also been
intensively studied, such as gas foaming and electrospinning.
In general, scaffolds for application in bone tissue engineering have to meet the
following requirements, 1) provide a biologically compatible and porous framework
to facilitate nutrients supply and waste removal, and 2) possess sufficient
mechanical strength to function from the time of cell seeding up to the point when
the host tissue begins to remodel the seeded construct after implantation [11].
More importantly, the cellular growth and osteogenic capacity of the seeded cells
should be supported/promoted by the scaffolds to form the ultimate bone tissue.
16
General Introduction
When the endochondral approach is applied, the scaffolds should also be able to
support cells in differentiating into the chondrogenic lineage, as the creation of an
in vitro cartilage template is necessary.
5. Objectives of this thesis
With the potential to overcome the drawbacks of the current autografts and
allografts, bone tissue engineering undoubtedly has been suggested as a promising
technique for reconstructing bone defects. However, until recently, the clinical
translation of this strategy is still unsatisfactory. Lack of timely and sufficient
vascular supply, resulting in immediate cell death after implantation, is generally
thought to be the cause of failure of the cell-scaffold constructs in patients [4]. To
overcome this issue, a potential strategy is to differentiate cells in a metabolically
less-demanding and angiogenic pathway, i.e. along the endochondral ossification
pathway, as this seems to hold promise for achieving better cell survival and
vascularization of the implanted constructs in vivo. Nevertheless, compared to
the available knowledge on bone regeneration via the intramembranous pathway,
such an endochondral approach is still in its infancy. Therefore, the overall
objective of thesis was to contribute to a better understanding of endochondral
bone tissue engineering. In particular, the studies focused on 1) exploiting new
porous scaffolds for bone regeneration applications, 2) establishing a cell-scaffold
model for studying endochondral bone formation, 3) testing the applicability
of endochondral approach with different scaffold materials, 4) investigating
the optimal in vitro priming time for the creation of cartilage templates, and 5)
exploring the underlying mechanism of the transformation from cartilage into
bone.
More specifically, the following research questions were addressed:
1. What is the current state of the art of combining stem cells and biodegradable
materials for bone regeneration via the endochondral pathway? (Chapter 2)
2. Is the porous aliphatic polyurethane/hydroxyapatite composite a suitable
scaffold for bone tissue engineering applications? (Chapter 3)
3. Is it possible to fabricate a 3D scaffold by electrospinning technique and utilize
it as cell carrier to enable effective in vivo endochondral bone generation?
(Chapter 4)
4. Can in vivo endochondral bone generation be achieved in various scaffold
materials and how is this process influenced by the scaffold properties?
(Chapter 5 and Chapter 6)
5. What is the effect of in vitro chondrogenic priming time of MSCs on in vivo
endochondral bone formation? (Chapter 6)
17
1
Chapter 1
6. What is the role of cartilage templates in mediating MSC behavior in the process
of endochondral bone formation? (Chapter 7)
References
1.
Giannoudis, P.V., Dinopoulos, H., and Tsiridis, E., Bone substitutes: an update. Injury, 2005. 36
Suppl 3: p. S20-7.
2. Buck, D.W., 2nd and Dumanian, G.A., Bone biology and physiology: Part II. Clinical correlates.
Plast Reconstr Surg, 2012. 129(6): p. 950e-956e.
3. Dawson, J.I. and Oreffo, R.O., Bridging the regeneration gap: stem cells, biomaterials and clinical
translation in bone tissue engineering. Arch Biochem Biophys, 2008. 473(2): p. 124-31.
4. Meijer, G.J., de Bruijn, J.D., Koole, R., and van Blitterswijk, C.A., Cell-based bone tissue
engineering. PLoS Med, 2007. 4(2): p. e9.
5. Langer, R. and Vacanti, J.P., Tissue engineering. Science, 1993. 260(5110): p. 920-6.
6. Bianco, P. and Robey, P.G., Stem cells in tissue engineering. Nature, 2001. 414(6859): p. 118-21.
7. Caplan, A.I., Review: mesenchymal stem cells: cell-based reconstructive therapy in orthopedics.
Tissue Eng, 2005. 11(7-8): p. 1198-211.
8. Gawlitta, D., Farrell, E., Malda, J., Creemers, L.B., Alblas, J., and Dhert, W.J., Modulating
endochondral ossification of multipotent stromal cells for bone regeneration. Tissue Eng Part B
Rev. 16(4): p. 385-95.
9. Farrell, E., van der Jagt, O.P., Koevoet, W., Kops, N., van Manen, C.J., Hellingman, C.A., Jahr, H.,
O’Brien, F.J., Verhaar, J.A., Weinans, H., and van Osch, G.J., Chondrogenic priming of human
bone marrow stromal cells: a better route to bone repair? Tissue Eng Part C Methods, 2009.
15(2): p. 285-95.
10. Scotti, C., Tonnarelli, B., Papadimitropoulos, A., Scherberich, A., Schaeren, S., Schauerte, A.,
Lopez-Rios, J., Zeller, R., Barbero, A., and Martin, I., Recapitulation of endochondral bone
formation using human adult mesenchymal stem cells as a paradigm for developmental
engineering. Proc Natl Acad Sci U S A, 2010. 107(16): p. 7251-6.
11. Hutmacher, D.W., Schantz, J.T., Lam, C.X., Tan, K.C., and Lim, T.C., State of the art and future
directions of scaffold-based bone engineering from a biomaterials perspective. J Tissue Eng
Regen Med, 2007. 1(4): p. 245-60.
12. Scaglione, S., Giannoni, P., Bianchini, P., Sandri, M., Marotta, R., Firpo, G., Valbusa, U., Tampieri,
A., Diaspro, A., Bianco, P., and Quarto, R., Order versus Disorder: in vivo bone formation within
osteoconductive scaffolds. Sci Rep, 2012. 2: p. 274.
18
General Introduction
19
1
Chapter 2
Combining osteochondral
stem cells and biodegradable
materials for bone regeneration
Wanxun Yang*, Matilde Bongio*, Fang Yang, Sanne K. Both,
Sander C.G. Leeuwenburgh, Jeroen J.J.P.
van den Beucken and John A. Jansen
*These authors contribute equally.
Combining Osteochondral Stem Cells and Biomaterials for Bone Regeneration
1. Introduction
Musculoskeletal malconditions (e.g. arthritis, back pain, spinal disorders,
osteoporosis, bone fractures, trauma, and congenital problems) affect hundreds of
millions of people across the world. According to ‘The Burden of Musculoskeletal
Diseases in the United States’ (BMUS), over 130 million patient visits to healthcare
providers occur annually, and the economic impact of these conditions is
staggering: in 2004, the annual costs for bone and joint health has been estimated
to be $849 billion dollars, and the burden of musculoskeletal conditions is expected
to escalate in the next 10-20 years as the population ages and lifestyles become
more active (http://www.boneandjointburden.org).
The most favorable choice for treatment of bone defects is the graft of the patient’s
own healthy bone (so-called autograft) harvested from a donor site, usually iliac
crest; or, alternatively, from a cadaveric bone typically sourced from a bone bank
(allograft). In addition to the respective benefits of these approaches, several
drawbacks limit their clinical applications. Autografts reduce the risk of rejection
but require an additional surgical site, thus increasing post-operative pain and
complications for the patient. Allograft eliminates the risk of donor site morbidity
and is hence associated with less pain and discomfort for the recipient; however,
rejection or disease transmission cannot be ruled out.
In view of these drawbacks associated with autografts and allografts, the
multidisciplinary field of regenerative medicine (RM) holds the promise of creating
appealing new possibilities regarding the quality and duration of life. Currently,
RM-based treatment strategies for critical-size bone defects (i.e. defects that will
not heal spontaneously) propose the design of biomaterials either alone or in
combination with stem cells and signaling molecules, which resemble the three
dimensional in vivo tissue arrangements and promote restoration of normal tissue
functions following grafting in the body [1]. To that end, a deep knowledge of bone
biology will lead to better understanding of the mechanisms of interaction between
cells and scaffolds, including cell adhesion, proliferation, and differentiation,
which are all important steps for the success of a man-made bone graft. Up to
now, thousands of scaffolds with variable properties and chemical composition,
including ceramics, natural and synthetic polymers, as well as cell culture
strategies have been proposed. Nevertheless, as for most of these approaches,
the developmental stage is in its infancy and the challenge is to obtain proof of
clinical efficacy in order to proceed towards routine clinical implementation.
The present chapter will firstly describe bone tissue and its developmental
mechanisms, in order to reach fundamental understanding of the processes
involved in RM-based approaches for bone regeneration utilizing biomaterials
23
2
Chapter 2
and cell culture strategies. Subsequently, current results in basic research on
biomaterials for bone repair, with emphasis on the endochondral ossification
approach, will be discussed.
2. Bone tissue
Bone is a dynamic and vascularized tissue with a remarkable ability to heal and
remodel, and it is responsible for providing structural support to the body and
protection of the internal organs. In addition, bone plays a key role in supporting
muscle contraction, blood production and the storage of mineral components [2].
Bone is not a uniformly solid structure, but is rather composed of two distinct
parts, namely compact and trabecular bone. Compact bone (also called cortical
bone) forms the hard outer component of the bones, whereas trabecular bone
(also called cancellous bone) fills the inner side. Compact bone contributes
approximately 80% of the weight of a human skeleton with a very dense and stiff
structure and is usually located in the shaft of the long bones and in the peripheral
layer of flat bones. In contrast, trabecular bone has a lower density, larger surface
area, and is composed of a highly porous network that creates space for bone
marrow and blood vessels.
From a material perspective, bone is composed of an inorganic and organic
component. The inorganic matrix represents ~70% of bone mass and is mainly
composed of hydroxyapatite, which provides stiffness and strength. The organic
matrix, which represents ~30% of bone mass, provides flexibility and tensile
strength, and is composed of proteins and cells. Collagen type I is the main
component of the organic matrix; whereas other proteins, such as osteocalcin,
bone sialoprotein and osteonectin, play important signaling functions and are
involved in mineralization process.
The three types of cells that populate bone tissue are osteoblasts, osteocytes and
osteoclasts. Osteoblasts are bone-forming cells that originate from osteoprogenitor
cells located in the deeper layer of periosteum and bone marrow, which produce
an osteoid matrix that successively mineralizes into bone. Although the term
osteoblast implies an immature cell type, osteoblasts are the mature cells of bone
that generate bone tissue. At a later stage, osteoblasts can either become bone
lining cells or differentiate into osteocytes. Bone lining cells are inactive osteoblasts
that reside along the quiescent bone surface forming a protective layer. Although
their function is not well understood, they may play a role in the activation of bone
remodeling [3]. Osteocytes are embedded within bone matrix in so-called bone
lacunae. These mature bone cells are involved in different functions, including
matrix maintenance, calcium homeostasis and regulation of bone response to
24
Combining Osteochondral Stem Cells and Biomaterials for Bone Regeneration
mechanical stimuli. Osteoclasts are responsible for bone resorption by (locally)
decreasing the pH and activating lytic enzymes, which in turn break down the
mineralized matrix.
2.1 Osteogenesis
During osteogenesis, two different processes can result in the formation of
normal, healthy bone tissue, i.e. intramembranous ossification and endochondral
ossification (Figure 1).
Intramembranous ossification
Intramembranous ossification occurs in a vascularized environment and involves
the direct differentiation of mesenchymal stem cells (MSCs) into osteoblasts.
Initially, MSCs are widely dispersed within an extracellular matrix that is devoid
of collagen. Groups of adjacent MSCs proliferate and form a dense nodule of cell
aggregation, which undergoes morphological changes toward the osteoblastic
phenotype. Subsequently, cells start to produce new organic matrix (i.e. osteoid),
mainly consisting of type I collagen fibrils, and eventually develop into mature
osteoblasts. Some osteoblasts align along the periphery of the nodule and keep
forming osteoid, whereas others are incorporated into the nodule and become
osteocytes. At this point, the osteoid becomes mineralized and forms the
rudimentary bone tissue.
Endochondral ossification
Endochondral ossification is the type of osteogenesis that accounts for the
development of most bones of the vertebrate skeleton (e.g. long bones) and is
the most commonly occurring process in fracture healing. As such, endochondral
ossification takes place in a poorly or non-vascularized environment, in which
loosely connected MSCs condense at the sites pre-determined for bone
formation. The increased cell-cell contacts and low oxygen condition (i.e. hypoxia)
trigger the differentiation of MSCs into the chondrogenic lineage. The cartilage
template takes shape through the deposition of abundant cartilaginous matrix,
and the chondrocytes gradually differentiate until reaching a hypertrophic stage,
during which the chondrocytes mineralize and eventually become apoptotic.
Simultaneously, these chondrogenic cells release vascular endothelial growth
factor (VEGF), which induces blood vessel formation [4]. At this point, osteoclasts
and osteoblasts, which receive nutrition and oxygen supply from the blood vessels,
reach the mineralized matrix and promote the transformation of the cartilage into
bone.
25
2
Chapter 2
Figure 1. Schematic representation of the two types of osteogenesis. 1)
Intramenbranous ossification involves bone formation through the direct osteoblastic
differentiation; 2) endochondral ossification generates bone via transition from an
avascular hypertrophic cartilage to a vascularized and mineralized matrix.
3. Basic tools for RM purposes - stem cells and biomaterial
scaffolds
3.1 Stem cells
Stem cells are undifferentiated cells with a high proliferation capability, which
have the capacity of self-renewal or differentiate into specialized cell types of our
body [5]. The central focus of RM are human stem cells, for which two main types
of stem cells exist based on their origin: embryonic stem cells (ESCs) and adult
stem cells. ESCs are derived from the inner cell mass of the blastocyst (i.e. an
early-stage embryo) and are able to differentiate into all cell types of the adult
organism. The multi-differentiation potential of ESCs has been reported by several
researchers [6, 7], and has attracted wide interest for many medical applications,
including regeneration of bone tissue. Indeed, ESCs have demonstrated the
capacity to differentiate into osteoblasts when cultured in the presence of
osteogenic differentiation-inducing factors, such as dexamethasone [8]. However,
many issues regarding the clinical use of ESCs need to be addressed, including
the risk of tumorogenesis, the labor-intensive procedures, and important ethical
26
Combining Osteochondral Stem Cells and Biomaterials for Bone Regeneration
considerations [9]. Adult stem cells, also called somatic stem cells, are present in
the body at many sites and are responsible for growth and maintenance of normal
tissues, as well as restoration of injured tissues. Initially, adult stem cells were
discovered in the bone marrow and, successively, in many other different tissues,
including heart, liver, pancreas, brain, epidermis, dental pulp and spinal cord. In
the last 30 years, substantial efforts have been devoted to studies of adult stem
cells for a broad range of applications for tissue regeneration. Among the different
sources for adult stem cells, mesenchymal stem cells (MSCs) are of great interests
in RM applications. MSCs have been initially identified in bone marrow as nonhematopoietic stem cells that can differentiate into tissues of mesodermal origin,
such as adipocytes, osteoblasts, chondrocytes, tenocytes, skeletal myocytes and
visceral stromal cells [10]. Friedenstein and co-workers were the first to reveal the
bone-forming capacity of MSCs [11].
3.2 Biomaterial scaffolds
As anticipated earlier in the introduction, a promising strategy to optimize
and simplify the treatment of muscoloskeletal disorders is represented by
bone grafts from man-made source. Several decades of progress in techniques
and understanding of the mechanisms involved in the interaction between
synthetic materials and living surrounding tissues resulted in the establishment
of ‘biomaterials’ as a scientific discipline, and a rapid progression in the design
of increasingly sophisticated materials that provide a template for tissue
regeneration. As a result of this evolution, the definition of “biomaterial” has
constantly changed over the years. Currently, the most appropriate and used
definition depicts biomaterial as “a substance that has been engineered to take a
form which, alone or as part of a complex system, is used to direct, by control of
interactions with components of living systems, the course of any therapeutic or
diagnostic procedure, in human or veterinary medicine” [12].
The history of biomaterials for bone repair is marked by three generations, namely,
from tolerant and bioinert (the first generation), to bioactive and osteoconductive
(the second generation), and to osteoinductive biomaterials (the third generation)
that induce neighboring tissue to become bone [13]. Primordial biomaterials were
meant to be tolerated by the body and to passively replace missing or damaged
tissue. The difficult adaptation of the human body to these foreign materials
raised the need to create improved artificial substitutes. These scaffolds reproduce
the complex three dimensional environment to which cells are exposed to and
actively participate in the dialog with the natural environment (i.e. bioactive). To
this end, bone healing requires scaffolds that provide a porous and biologically
compatible framework onto which (precursors of) bone-forming cells migrate
27
2
Chapter 2
(i.e. osteoconduction) and form new bone tissue, or even direct differentiation of
stem cells into mature osteogenic cells (i.e. osteoinduction) [14]. Last, but not least,
biomaterials should degrade over time in concert with the growth of new bone,
and the degradation products should be non-toxic and completely eliminated from
the body. Table 1 lists the most important properties and respective definitions for
appropriate bone graft materials.
Table 1. Properties and definitions for a bone graft material
Properties
Biocompatibility
Bioactivity
Definitions
Behavior of a biomaterial and its byproducts to elicit little or no immune
response into the body
Ability of a biomaterial to interact with the natural environment of the
body and determine a biological effect
The full process of obtaining firm anchoring and incorporation of a
Osteointegration
biomaterial into living bone without fibrous tissue formation at the
interface
Osteoconductivity
The ability of a biomaterial to stimulate the migration of adjacent boneforming cells in order to induce the growth of bone tissue
The ability of a biomaterial to stimulate the migration of undifferentiated
Osteoinductivity
cells and induce their differentiation into active osteoblasts, in order to
promote de novo bone formation
Biodegradability
The ability of a biomaterial to degrade in concert with the natural bone
remodeling process
3.2.1 Ceramic-based materials
In the past three decades, ceramics, such as hydroxyapatite (HA), tricalcium
phosphate (TCP) and bioglass, were the main scaffolds studied for bone
regeneration, being similar in their chemical composition and crystallographic
structure to the mineral component of bone tissue. These characteristics govern
the bioactive and osteoconductive behavior, which in turn favor the formation of
new bone; although the brittleness and weak mechanical properties represent the
major drawbacks that still require optimization [15].
3.2.2 Polymer-based materials
An alternative approach in RM is to engineer new bone tissues by using selective
cell transplantation on polymer scaffolds [16]. Some of the most commonly explored
polymers as scaffold materials include poly(lactic acid) (PLA), poly(lactic-co-glycolic
28
Combining Osteochondral Stem Cells and Biomaterials for Bone Regeneration
acid) (PLGA), poly(2-hydroxyethyl methacrylate) (pHEMA), poly(ε-caprolactone)
(PCL), poly(ethylene glycol) (PEG), polyglycolic acid (PGA), polyurethanes and
composites [17]. Polymers exhibit biocompatibility, degradability and have good
processability. These properties confer to polymers a great potential to serve
as artificial ECM, thus mimicking and recreating in some extent the structural
organization of bone.
4. Cell culture strategies for bone tissue regeneration
Initial in vitro culture set-up, including selection of cell types (mono- or coculture), growth factors, medium composition and timing are pivotal to determine
the possible benefits of well-timed implantation [4]. Figure 2 depicts all the steps
required for the synthesis of a cell-scaffold construct for bone tissue regeneration,
which include (1) cell isolation and expansion in complete medium, (2) cell seeding
and cell maturation (via intramembranous or endochondral ossification), either
prior to or after loading into 3D scaffolds, and (3) in vivo grafting.
4.1 Isolation and expansion
As described above, natural bone regeneration is a complex process, during which
multiple cell types, dispersed in a three dimensional environment, are spatially
and temporally regulated by several environmental stimuli. In order to re-create
the natural situation, millions of cells are necessary for incorporation into an
artificial matrix. From a clinical point of view, it is possible to isolate osteoblasts
from biopsies taken from the patients, however, this procedure is time consuming
and allows to obtain only a limited number of cells. Therefore, stem cells represent
a favorable alternative. Stem cells, derived from bone marrow or other tissues, are
usually expanded with complete medium in culture plates or flasks, where they
adhere, start proliferating and form colony-forming cell clusters. Before these cells
become completely confluent, they are passaged and expanded in multiple flasks,
thus becoming a more and more homogenous adherent cell population that may
proliferate without differentiating up to several generations [18].
4.2 Seeding and differentiation
Intramembranous ossification pathway
Except for a few reports, regenerative medicine efforts in the field of bone tissue
repair have focused on generating bone directly from osteoprogenitor cells in a
manner akin to intramembranous ossification in vivo. The standard procedure of
this strategy involves stem cell differentiation into osteoblasts. To that end, MSCs
are exposed to specific culture supplements, such as dexamethasone, ascorbic
29
2
Chapter 2
acid, and beta glycerolphosphate [19]. Cells can either be differentiated in flasks
(2D-environment) and loaded into scaffolds just prior to grafting in vivo, or be
loaded into scaffolds when they are still uncommitted and undergo maturation
in a 3D-environment. Although the second approach is more appealing, as it
resembles the in vivo situation, its success has not been satisfactory to date. A
major issue with this approach is the limited long-term cell survival and necrosis,
arising from lack of nutrient delivery to and waste removal from the center of
tissue-engineered constructs, and lack of blood supply [20]. First, blood vessels
formation is not an instantaneous process and usually occurs too slowly for cells
to survive without nutrients. Second, culturing a cell-laden scaffold in vitro for
several weeks can lead to the formation of an extensive extracellular matrix, which
can prevent or seriously hamper blood vessel infiltration in vivo by sealing the
pores of a scaffold [21].
Endochondral ossification pathway
Because chondrocytes can survive in a hypoxic environment and they produce
anabolic and catabolic factors important for the conversion of avascular tissue
into vascularized tissue, it seems straightforward to exploit the chondrocytemediated route (i.e. endochondral) for ossification in RM-based approaches,
rather than intramembranous ossification [4]. Either human ESCs and MSCs are
able to form bone via the endochondral pathway, i.e. differentiate to chondrocytes
in vitro and turn into bone when implanted in vivo [20, 22]. Chondrogenic priming
of mesenchymal stem cells can be obtained by culturing cells in differentiation
medium containing a mixture of specific supplements, such as dexamethasone,
ascorbic acid, insulin, transferrin, selenious acid, sodium pyruvate, transforming
growth factor-beta (TGF-β), and a phosphate donor [23]. Three-dimensional
conformation of the cells in aggregates with high cell density and cell-cell
interaction play an important role in the mechanism of cartilage formation. The
main idea would be to induce the secretion of the angiogenic factors from the
hypertrophic chondrocytes or in vitro pre-vascularization of the templates in order
to facilitate blood vessel formation when implanted in vivo. However, no in vitro
system is known to support vascularization necessary for the formation of bone
following the endochondral ossification pathway. Different culture systems can
offer a solution for the development of in vitro tissue engineering approaches
for the creation of vascularized bone in vivo. Among these strategies, the use of
bioreactors [24] and an environment with low oxygen tension have demonstrated to
support growth and differentiation of cartilage that forms bone [25].
30
Combining Osteochondral Stem Cells and Biomaterials for Bone Regeneration
2
Figure 2. Illustration of all the fundamental steps required for the synthesis and
application of a cell-material system for bone tissue regeneration. (1) cell isolation
and expansion, (2) cell seeding and differentiation, and (3) in vivo grafting.
31
Chapter 2
4.3 In vivo grafting
Osteo- or chondroprogenitor cells, previously differentiated from stem cells under
prolonged conventional culture conditions, are implanted in vivo in conjunction
with natural or synthetic scaffolds. Osteoprogenitor cells differentiate directly into
osteoblasts that will deposit new bone matrix, while chondrogenic-committed
cells fulfill hypertrophy and further deposit cartilaginous extracellular matrix that
becomes calcified, and eventually undergo apoptosis. Meanwhile, blood vessels
and osteoblasts differentiated from the surrounding stem/progenitor cells, invade
the hypertrophic cartilage and give rise to bone [26].
5. Preliminary investigations on cell-biomaterial constructs for
endochondral bone formation
Regarding endochondral ossification, it is a very unique transition process,
which includes two distinguished stages: chondrogenesis and osteogenesis.
Cartilage matrix consists of highly hydrated proteoglycan embedded into a type
II collagen network. On the other hand, bone tissue is more rigid than cartilage
and impregnated mainly with hydroxyapatite (HA) and type I collagen. Regarding
the endochondral bone tissue engineering, the scaffolds should be able to
support cells to form transient cartilage-like tissue in vitro and also promote the
calcification of such cartilage template into bone after in vivo implantation. The
following sections describe the cell-scaffold systems that have been explored so
far for the generation of bone tissue via endochondral ossification approach.
5.1 Calcium phosphate (CaP) ceramics
Before being used as scaffold materials, CaP ceramics have been widely applied as
bone substitutes and implant coatings because of their resemblance to the mineral
composition of the natural bone. Commercially available CaP-based biomaterials
include HA, β-tricalcium phosphate (β-TCP), and biphasic CaP (BCP, containing
HA and β-TCP). The CaP ceramics are generally considered to be bioactive and
osteoconductive, which posses the ability to guide formation of the new bone tissue
tightly along their surfaces. This has made them the superior scaffold candidates
for bone tissue engineering. However, the development of three-dimensional in
vitro models of bone using bioceramics remains an important challenge that is
still to be overcome. CaP ceramics are reactive and their reactivity depends on
their characteristics (such as composition, dissolution, sintering temperature,
microstructure). When ceramics are cultured in vitro, calcium and phosphates are
released in the medium and, as a result, can hamper cellular function [27]. Since
testing the in vitro cellular activity is a pertinent step, and scaffolds and cells are in
32
Combining Osteochondral Stem Cells and Biomaterials for Bone Regeneration
contact for a certain period in vitro before being implanted, a number of studies
have been focusing on solving the difficulties of removal of excessive ions, as well
as oxygen and nutrients supply to cells within ceramic scaffolds [28]. One solution
is using bioreactors to convey the medium directly throughout the interconnected
pores to continuously introduce nutrients and remove wastes, so that the cell
viability and activities can be favored in long-term culture within 3D ceramic
scaffolds [29].
Teixeira et al. reported that a BCP scaffold, consisting of 60% HA and 40% β-TCP,
had the potential to support the chondrocytes isolated from chick embryonic
tibia to form bone via the endochondral route [15]. Tortelli et al. compared the
biological responses between murine MSCs and mature osteoblasts, when seeded
in HA scaffolds and subsequently implanted in immunocompromised mice. The
results showed that new vascularized bone was formed through the activation of
an endochondral ossification process in the MSCs-scaffolds complex. Conversely,
osteoblasts directly formed bone via an intramembranous ossification, without
however any signs of vascularization [30]. In another study, mouse embryonic stem
cells (ESCs) seeded onto ceramic scaffolds and cultured in osteogenic medium,
failed to form bone tissue upon implantation in subcutaneous pockets of nude
mice. Differently, when ESCs-ceramic constructs were cultured in chondrogenic
differentiation medium, they showed cartilage maturation toward the hypertrophic
stage, calcification and ultimately bone formation in vivo [22]. All together, these
studies indicated that the origin of bone cells and the ossification type are closely
related to the nature and commitment of the seeded cells.
5.2 Hydrogels
Hydrogels are networks of hydrophilic, crosslinked polymer chains with a distinctive
quality to absorb water in aqueous fluids, and hence to exhibit mechanical and
visco-elastic properties that closely resemble the structural three dimensional
environment of natural extracellular matrix (ECM). These attributes make
hydrogels appealing substrates for cells to adhere, proliferate, secrete new ECM
and eventually restore the damage tissue [31]. Most importantly, from a clinical
perspective, hydrogels are available ready-to-use grafts, can be injected into the
wound site using minimally invasive techniques and allow complete filling of
irregularly shaped defects, thus avoiding an open surgery procedure. In combination
with stem cells, these scaffolds have been evaluated for their suitability as potential
replacements for bone and cartilage. Hydrogels can be broadly classified into
natural and synthetic, based on their origin [32]. Natural hydrogels (e.g. agarose,
alginate, chitosan, collagen/gelatin, fibrin, hyaluronan, and silk), exhibit excellent
bioactivity, as they mimic the native extracellular environment structurally, have
33
2
Chapter 2
unique mechanical properties and are biodegradable by enzymatic or hydrolytic
mechanisms. Synthetic hydrogels can be formed by crosslinking of polymer chains
through physical (temperature, UV light or pH) or chemical reactions under
controlled conditions. Representative synthetically derived polymers explored in
combination with stem cells for bone regeneration purposes include poly(ethylene
glycol) (PEG), poly(vinyl alcohol) (PVA), oligo(poly(ethylene glycol) fumarate) (OPF)
hydrogels, and poly(ethylene glycol)-diacrylate (PEG-DA) hydrogels [33-35].
In regenerative medicine and tissue engineering of bone, the vast majority of
studies have been focused on the formation of new bone directly from stem cells
or mature osteogenic cells encapsulated within hydrogel matrices. However, more
recently, stimulation of cells toward the endochondral ossification using cell-laden
hydrogel systems has shown to be not only appealing for the understanding of
the regulatory mechanisms of cartilage and bone formation, but also a potential
approach for bone repair.
5.2.1 Alginate
Alginate is a linear heteropolysaccharides composed of D-mannuronic acid and
L-guluronic acid, and are derived primarily from brown algae. Due to the presence
of the carboxyl groups along the polymer chain, alginates can form gels in the
presence of divalent ions such as calcium ions [36]. Due to easy preparation under
gentle conditions, low toxicity and readily availability, alginate hydrogels have
been extensively investigated as matrices for encapsulating cells and delivering
drugs. Alginate hydrogels have been shown to be suitable matrices for generation
of bone tissue via endochondral ossification when associated with committed
stem cells and appropriate combinations of regulatory signals (e.g. BMP-2, VEGF,
TGF-β3) [37]. Chang et al. followed the progression of osteogenically induced MSCs
along the endochondral pathway [38]. At 2 weeks after subcutaneous implantation,
cell-alginate constructs presented islands of cartilage formation and cells had
the typical round morphology of the chondrocyte phenotype. By 6 weeks,
endochondral ossification with trabecular bone deposition started to appear,
whereas osteoblasts and osteocytes were seen in the new bone at week 8 and at
week 12, respectively. Simultaneously, increasing calcification was detected over
time. Similarly, in another study, hMSC were suspended in alginate beads with a
chondrogenesis-induction medium containing transforming growth factor TGF-β3.
During the first stage of culture, specific chondrogenic markers, such as collagen
type II, type X and proteoglycan, were detectable, and their expression decreased
over time in concert with the increase of osteogenic markers, such as osteocalcin
[39]
.
34
Combining Osteochondral Stem Cells and Biomaterials for Bone Regeneration
5.2.2 Chitosan
Chitosan is a linear polymer of (1-4) β-linked d-glucosamine residues with N-acetyl
glucosamine groups derived from chitin, a naturally occurring polysaccharide which
forms the outer shell of crustaceans, insect exoskeletons, and fungal cell walls.
Chitosan is completely soluble in aqueous solutions with pH lower than 5.0 and it
undergoes biodegradation in vivo enzymatically by lysozyme to nontoxic products
[40]
. Biocompatibility, biodegradability and structural similarity to natural GAGs
make crosslinked chitosans hydrogels attractive materials for tissue engineering
applications. Moreover, the feasibility of forming porous scaffolds showed to
promote osteoblastic differentiation in vitro and might support angiogenesis and
osteogenesis in vivo [41]. Oliveira et al. fabricated the chitosan sponges via freezedrying chitosan acidic solutions and further demonstrated that these chitosan
sponges with adequate pore structure and mechanical properties could serve as
a support for hypertrophic chondrocytes to induce endochondral ossification [42].
5.2.3 Collagen
Collagen represents the most abundant protein in the human body, being the
major component of bone, skin, ligament, cartilage, and tendon. It also forms
the structural framework for other tissues such as blood vessels. There are
at least 19 different types of collagen, but the basic unit of all is a polypeptide
consisting of a three-amino-acid sequence (glycine, proline, and hydroxyproline)
forming polypeptide chains, which wrap around one another to form the lefthanded triple helix structure [43]. Due to its excellent biocompatibility, enzymatic
degradability and resemblance to the organic composition of natural bone,
collagen-based matrices find wide application in tissue engineering and have
been extensively investigated in combination with cells and growth factors [44].
Abrahamsson et al. seeded human MSCs suspended collagen gel on a non-woven
PCL scaffold. Cultured in chondrogenic condition, cartilaginous tissue was formed
by day 21, and hypertrophic mineralization was observed in the newly formed
ECM at the interface with underlying scaffold by day 45 [37]. Collagen hydrogels
could also be fabricated into porous sponges before cell seeding. MSCs-collagen/
glycosaminoglycan constructs pre-cultured in chondrogenic culture medium and
subsequently implanted subcutaneously in the dorsum of nude mice showed
good cell survival and progression along the endochondral ossification route,
and more interesting, blood vessels formation. Conversely, the osteogenically
primed scaffolds showed a mineralized matrix of poor quality, no vascularization,
likely due to too much matrix deposition or a lack of release of inductive factors,
and few surviving cells, as a result of absence of oxygen and nutrients [4, 20]. In
another study, chondrocytes isolated from chick embryo were cultured in
collagen sponges treated with retinoic acid to induce chondrocyte maturation
35
2
Chapter 2
and extracellular matrix deposition. Biological properties and stiffness of collagen
matrices supported chondrocytes attachment, proliferation and endochondral
bone maturation after implantation [45].
6. Conclusion and future perspective
Regenerative medicine is a multidisciplinary field aiming at restoration of normal
function to injured tissues, by the development of increasingly sophisticated
biomaterials, either alone or in combination with cells and/or stimulating factors.
The enormous progress achieved so far at the laboratory stage and the continuous
worldwide effort towards new insights and technologies, direct researchers to face
new challenges. In order to translate basic scientific advantages into approved
medical products, all the materials requirements (e.g. safety and efficacy) must
comply the guidelines established by regulatory agencies, including the Food and
Drug Administration Agency (FDA) and the European Medicinal Agency (EMEA),
and clinical trials must be carried out within acceptable clinical practice and due
ethical considerations.
In conclusion, although great challenges still remain, research on combining
stem cells with biomaterial for the creation of bone grafts is at the precipice of
major breakthroughs that likely will change and simplify the treatment of bone
disorders, thus to bring major benefits to patients, including reduced need and
cost of medical attention, and short-term recovery.
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38
Chapter 3
Biological evaluation of
porous aliphatic polyurethane/
hydroxyapatite composite scaffolds
for bone tissue engineering
Wanxun Yang, Sanne K. Both, Yi Zuo, Zeinab Tahmasebi Birgani,
Pamela Habibovic, Yubao Li, John A. Jansen and Fang Yang
Biological Evaluation of PU/HA scaffolds
Summary of this chapter
Biomaterial scaffolds meant to function as supporting structures to osteogenic cells
play a pivotal role in bone tissue engineering. Recently, we synthesized an aliphatic
polyurethane (PU) scaffold via a foaming method using non-toxic components.
Through this procedure a uniform interconnected porous structure was created.
Furthermore, hydroxyapatite (HA) particles were introduced into this process to
increase the bioactivity of the PU matrix. To evaluate the biological performances
of these PU based scaffolds, their influence on in vitro cellular behavior and in vivo
bone forming capacity of the engineered cell-scaffold constructs was investigated in
this chapter. A simulated body fluid (SBF) test demonstrated that the incorporation
of 40wt% HA particles significantly promoted the biomineralization ability of the
PU scaffolds. Enhanced in vitro proliferation and osteogenic differentiation of
the seeded mesenchymal stem cells (MSCs) were also observed on the PU/HA
composite. Next, the cell-scaffold constructs were implanted subcutaneously in
a nude mice model. After 8 weeks, a considerable amount of vascularized bone
tissue with initial marrow stroma development was generated in both PU and PU/
HA40 scaffold. In conclusion, the PU/HA composite is a potential scaffold for bone
regeneration applications.
1. Introduction
Since the concept of tissue engineering was introduced, researchers have been
seeking a solution for bone repair and regeneration by combining the three
key factors, i.e. a porous scaffold, exogenous stem cells and biological cues.[1]
Essentially, the choice of a biomaterial as scaffold is critical to the success of bone
tissue engineering.[2] It is well accepted that an optimal scaffold should possess a
porous and biologically compatible framework onto which the bone-forming cells
can attach, function, and eventually form new bone tissue.[3]
Recently, polyurethanes (PUs) gained popularity and have been investigated as
scaffold material. A typical PU is a block copolymer consisting of a hard segment
contributed by isocyanates and a soft segment formed by polyether or polyester
polyols. For that reason, the material properties of PU, such as mechanical strength,
biodegradability and cytocompatibility, can be easily modified by adjusting the
components of the hard and soft segments during PU synthesis.[4, 5] For instance,
a PU synthesized from aliphatic diisocyanate, e.g. isophorone diisocyanate (IPDI),
non-toxic castor oil and 1,4-butanediol (BDO) gives rise to non-toxic degradation
products, which is more desirable to be used as a scaffold material than the
conventional PUs made from aromatic diisocyanates.[6, 7] However, like most
43
3
Chapter 3
synthetic polymers, PUs lack bioactive groups to facilitate biomineralization.
The most common strategy to counteract the poor bioactivity is by introducing
bioactive ceramics, e.g. hydroxyapatite (HA) particles, into the PU matrix during
the polymerization process.[8, 9] Due to the presence of the polar groups in the
molecular chain, PUs have relatively high affinity to HA.[10]
Another attractive feature of PUs is that they have a preference to spontaneously
foam during the copolymerization process, which facilitates a one-step process of
making porous scaffolds. For instance, water, which is either unavoidably present
in the reaction mixture or added intentionally during the copolymerization
process, reacts with the isocyanate causing the release of CO2 gas. The released
CO2 gas creates bubbles and eventually leads to foaming. Our previous study
has demonstrated that a mild foaming process could be achieved to allow the
formation of a uniform porous PU structure by carefully choosing a low reaction
rate isocyanate.[11] We also demonstrated that nanosized HA particles could be
added into this process, resulting the formation of a porous PU/HA composite
scaffold with good dispersion and high occupancy of HA particles.[11, 12]
In this study, we aimed to further test the biological performance of such PU/
HA scaffolds in comparison to HA free PU scaffolds, including the scaffold effect
on cellular behavior and in vivo bone forming capacity of the engineered cellscaffold constructs. Firstly, the scaffolds were immersed in simulated body fluid
(SBF) to test biomineralization ability. Furthermore, rat bone marrow derived
mesenchymal stem cells (MSCs) were seeded onto the scaffolds, and the cell
viability and osteogenic capacity were tested in vitro. Finally, the efficacy of bone
formation on the cell-scaffold constructs was evaluated in vivo after subcutaneous
implantation in nude mice.
2. Materials and methods
2.1 Scaffold fabrication and morphology characterization
All the reagents were of analytic grade and purchased from Kelong Co. Ltd.
(Chengdu, China), except that castor oil and isophorone diisocyanate (IPDI)
were obtained from Aladdin Co. Ltd. (Shanghai, China). Both castor oil and
1,4-butanediol (BDO) were dehydrated under decompression with a vacuum of
1330 Pa at 120°C. IPDI was preserved in a refrigerator before use. Nano-sized HA
slurry was prepared by wet synthesis as previously described,[13] then spray-dried
to particles of approximately 5-15 μm in diameter for further experiments.
The PU and PU/HA scaffolds were prepared by copolymerization and simultaneous
foaming following our previously reported method [11]. The reaction was performed
44
Biological Evaluation of PU/HA scaffolds
in a 250 ml three-necked round-bottom flask under a dry nitrogen atmosphere.
Firstly, 38 g castor oil, and a certain amount of HA powders, were put into the
flask and stirred uniformly. Subsequently, 11.2 g IPDI was added drop-wise into the
HA/castor oil mixture and the reaction was kept for 4 h to form the prepolymer.
Then 4.5 g BDO was used as a chain extender to cross-link the prepolymer and
0.9 g deionized water was added in the cross-linked prepolymer by stirring for
5 min. The mixture was placed in an oven at 120°C for 4 h accompanied with
simultaneous foaming, after which the three-dimensional porous bulk scaffold
was obtained. Three types of scaffolds were prepared based on the amount of HA.
They were named as PU, PU/HA20 and PU/HA40 with the PU:HA weight ratios of
100:0, 80:20 and 60:40, respectively. For all three types of scaffolds, the overall
porosity was 78-81%.[11]
Disk-shaped samples with a diameter of 6 mm and a thickness of 2.5 mm were
punched out of the bulk scaffolds. In order to achieve further exposure of the
HA particles on the surfaces, all types of scaffolds were immersed in 20M NaOH
solution for 5 days under gentle agitation, as NaOH has the erosive effect on PU.[14]
After that all scaffolds were cleaned thoroughly in deionized water and freezedried before being used in the further experiments. The scaffold morphology was
examined by scanning electron microscopy (SEM; JEOL6340F, Tokyo, Japan) after
being sputter-coated with gold.
2.2 Simulated body fluid (SBF) immersion test
SBF immersion test was used to evaluate the biomineralization behavior of the
scaffolds. The recipe of SBF was adopted from Kokubo [15] and the pH value of
the solution was adjusted to 7.4 after complete preparation. Immersion studies
were performed by incubating each sample in 10 ml SBF solution in a 15 ml tube
(n=3). All tubes were placed in a water bath at 37 °C under continuous shaking.
After immersion periods of 1 and 4 weeks, the scaffolds were gently washed with
deionized water and freeze-dried. SEM and elemental analysis of the deposits
was carried out using a Philips XL30 scanning electron microscope (Eindhoven,
the Netherlands) equipped with an energy dispersive spectrometer (EDS, AMETEK
Materials Analysis Division, Mahwah, NJ, USA) after the samples were sputtercoated with gold (Cressington 108A, Watford, UK). The accelerating voltage was
10 KeV and the working distance was 10 mm. EDS analysis provided information
on the distribution of elements of interest (Ca and P) in the analyzed area. 2.3 Cell isolation and seeding
Rat MSCs were isolated from 6-week-old male Wistar rats after the approval from
Radboud University Nijmegen Animal Ethics Committee. Briefly, two femora of
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each rat were extracted and the epiphyses were cut off. MSCs were flushed out
of the remaining diaphyses using the primary cell culture medium, consisting of
alpha Minimal Essential Medium (aMEM; Gibco, Grand Island, USA) supplemented
with 10% fetal bovine serum (FBS; Sigma, St. Louis, USA), 100 U/ml penicillin and
100 μg/ml streptomycin (Gibco). The flush-out was cultured for two days in a
humidified incubator (37 °C, 5% CO2), after which the medium was refreshed to
remove non-adherent cells. Prior to detaching the cells, they were cultured for an
additional three days. Then, the cells were detached using trypsin/EDTA (0.25%
w/v trypsin, 0.02% w/v EDTA; Sigma) and counted.
The PU scaffolds for cell culture experiments were sterilized by autoclave and prewetted in primary cell culture medium overnight. For cell seeding, 5× 104 cells
were suspended in 50 µl medium and statically seeded onto each scaffold. All
scaffolds were placed in 24-well non-adherent culture plates (1 scaffold per well)
and incubated for 3 h in a humidified incubator (37 °C, 5% CO2) allowing the initial
cell attachment to the scaffolds. Subsequently, osteogenic medium was added,
which contained 50 μg/ml ascorbic acid (Sigma), 10nM dexamethasone (Sigma)
and 10 mM sodium β-glycerophosphate (Sigma) on the basis of the primary
culture medium. Scaffolds were maintained in culture for 24 days, with the cell
culture medium being refreshed twice per week. Cell viability, DNA amount,
alkaline phosphatase (ALP) activity and osteogenic gene expression were tested.
2.4 Cell viability
LIVE/DEAD® Assay (Invitrogen) was used to assess cell viability after 24 h of cell
seeding. The cell-scaffold constructs were washed in PBS and exposed to calcein
AM / ethidium homodimer-1 / PBS working solution for 30 min at 37 °C according to
the manufacturer’s instructions. Dye uptake was detected by using an automated
fluorescent microscope (Axio Imager Microscope Z1; Carl Zeiss Micro Imaging
GmbH) with a wave length of 488 nm for visualizing the live cells (green) and 568
nm for the dead cells (red).
2.5 DNA content and ALP activity
DNA content (n=3) was quantified by PicoGreen assay (Quant-iT PicoGreen dsDNA,
Invitrogen) after 4, 8, 16 and 24 days of osteogenic culture condition. After the
culture medium was removed, scaffolds were washed twice with PBS. Cells were
harvested by placing the cell-scaffold constructs in a 1.5-ml tube. One ml of
deionized water was added to each sample, after which 2 freeze-thaw cycles and
ultrasonication were performed. One-hundred μl cell lysate or DNA standard from
the PicoGreen assay kit was added to 100 μl freshly made working solution in
the wells. The results were read using a fluorescence microplate reader (Bio-Tek
46
Biological Evaluation of PU/HA scaffolds
Instruments, Abcoude, The Netherlands) with an excitation wavelength at 485nm
and an emission wavelength at 530nm.
The ALP activity (n=3) was measured as a marker for early osteogenic differentiation
using the same samples as for the DNA assay. Eighty μl of cell lysate or standard
(serial dilutions of 4-nitrophenol at the concentrations of 0-25 nM) and 20μl of
buffer solution (5 mM MgCl2, 0.5 M 2-amino-2-methyl-1-propanol) were added into
a 96-well plate. Then 100μl of substrate solution (5 mM paranitrophenylphosphate)
was added into all wells. Subsequently, the plate was incubated for 1 h at 37 °C
before the reaction was stopped by adding 100 μl of 0.3M NaOH. The plate was
read in an ELISA reader (Bio-Tek Instruments) at 405 nm. ALP activity results were
normalized by the amount of DNA.
2.6 RNA extraction and real-time PCR analysis
To analyze the expression of osteogenic-related genes of the cells seeded on each
scaffold, a real-time PCR was performed. Total RNA was extracted using Trizol®
method (Invitrogen) after 4 and 8 days of culturing. Briefly, scaffolds with cells were
washed with PBS before 1 ml of Trizol® solution was added. The cell extract was
then collected, mixed with chloroform and centrifuged. Only the upper aqueous
phase was collected and mixed with equal amount of isopropanol. After 10min of
incubation at room temperature, the extract was centrifuged and washed twice
with 75% alcohol. The RNA pellet was dissolved in RNA free water and the total
RNA concentration was measured with spectrophotometer (NanoDrop 2000,
Thermal scientific, Wilmington, USA).
Table 1. Overview of the primer sequences
Forward (5’ → 3’)
Reverse (5’ → 3’)
Oc
CGGCCCTGAGTCTGACAAA
GCCGGAGTCTGTTCACTACCTT
Bsp
TCCTCCTCTGAAACGGTTTCC
GGAACTATCGCCGTCTCCATT
Runx-2
GAGCACAAACATGGCTGAGA
TGGAGATGTTGCTCTGTTCG
Gapdh
CTTCACCACCATGGAGAAGGC
GGCATGGACTGTGGTCATGAG
First strand cDNA was reverse transcribed from RNA using the SuperScript FirstStrand Synthesis System kit (Invitrogen). Afterwards cDNA was further amplified
and the expression of specific genes was quantified using IQ SYBR Green Supermix
PCR kit (BioRad, Hemel Hempstead, United Kingdom) in a real-time PCR (BioRaD,
CFX96™ real-time system). Osteogenic markers expressed on RNA level were
evaluated, including osteocalcin (Oc), bone sialoprotein (Bsp) and runt-related
transcription factor 2 (Runx2) (Table 1). The expression levels were analyzed and
47
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Chapter 3
compared to the housekeeping gene glyceraldehyde-3-phosphate dehydrogenase
(Gapdh). The specificity of the primers was tested separately before the real-time
PCR reaction. The expression of the tested genes was calculated via the 2-∆∆Ct
method [15] and the PU scaffolds were used as the reference group.
2.7 In vivo implantation
Based on the in vitro results, two types of scaffolds, PU and PU/HA40 with
distinct biological properties were sterilized by autoclave and selected for the in
vivo study. The protocol was approved by the Animal Ethical Committee of the
Radboud University Nijmegen Medical Centre (Approval no: RU-DEC 2010-254)
and the national guidelines for the care and use of laboratory animals were
applied. Before in vivo implantation, 250,000 cells were seeded on each scaffold
and cultured for 5 days in vitro. Afterwards, the cell-scaffold constructs were
implanted subcutaneously in three male nude mice (HsdCpb:NMRI-nu, Harlan).
Surgery was performed under general inhalation anesthesia with a combination
of isoflurane, nitrous oxide, and oxygen. All mice received analgesic before and
after the surgery. The back of the mice was shaved and disinfected with povidoneiodine. Subsequently, four small longitudinal incisions were made. Lateral to
each incision, a subcutaneous pocket was created. A cell-scaffold construct and
a cell-free scaffold from both PU and PU/HA40 groups were implanted in each
mouse. Afterwards the skin was closed using staples. After 8 weeks, the mice were
euthanized by CO2-suffocation for sample collection.
2.8 Histological analysis
All the in vivo samples were fixed in 10% phosphate buffered formalin for 30 h
and dehydrated through graded ethanol before embedding in methylmetacrylate
(MMA, L.T.I., Bilthoven, the Netherlands). Embedded samples were sectioned using
a microtome equipped with diamond blade (Leica SP1600, Leica microsystem,
Wetzlar, Germany) and stained with methylene blue and basic fuchsin (both
reagents from Merck). The staining was imaged using a light microscope (Zeiss
Imager Z1, Carl Zeiss AG Microscopy, Germany) equipped with AxioCam MRc5
camera.
2.9 Statistic analysis
Statistical significance in this study was evaluated using ANOVA analysis followed
by Tukey’s Multiple Comparison Test (GraphPad Software Inc., San Diego, USA).
Results are reported as mean values and standard deviation. Differences were
considered statistically significant at p < 0.05.
48
Biological Evaluation of PU/HA scaffolds
3. Results
3.1 Scaffold morphology
As shown in the SEM micrographs (Figure 1), the PU and PU/HA scaffolds displayed
a porous structure, containing both macropores ranging from 300 to 1000 μm
and micropores from 50 to 300 μm which were mainly located on the surface
of macropores (Figure 1a-c). Compared to PU scaffolds, PU/HA20 and PU/HA40
scaffolds exhibited a less regular shape of the macropores and also a less number
of the micropores. In addition, due to the presence of the HA particles, the two
PU/HA composites showed an unevenness of the pore wall topography compared
to the PU scaffolds (Figure 1d-f).
Figure 1. SEM micrographs of the PU(a and d), PU/HA20 (b and e) and PU/HA40 (c
and f) scaffolds. The PU and PU/HA scaffolds displayed a porous structure, containing
both macropores and micropores. Compared to PU scaffolds, PU/HA20 and PU/HA40
scaffolds exhibited a less regular shape of the macropores and also a less number of
the micropores (a-c). In the magnified figures (d-f), an increased unevenness on pore
surfaces of PU/HA20 and PU/HA40 was observed due to the presence of HA particles.
(White arrows indicate the micropores.)
3.2 Biomineralization test
After 1-week immersion in SBF solution, no obvious calcium phosphate (CaP)
precipitation was observed on either PU or PU/HA20 scaffolds by SEM, whereas
EDS detected marginal quantities of calcium and phosphorus (Figure 2a-d). In
comparison, a significant extent of coverage by flake-like crystal structure was
49
3
Chapter 3
Figure 2. Biomineralization of the scaffolds after 1-week immersion in SBF. EDS
detected marginal quantities of calcium and phosphorus on PU (a) and PU/HA20 (c)
scaffolds, while no obvious CaP precipitation was observed on these scaffolds by SEM
(b and d). In comparison, significant coverage by flake-like crystal structure was
observed on the PU/HA40 scaffold (e), from which a considerable amount of calcium
and phosphorus were detected by EDS (f).
observed on the surfaces of the PU/HA40 scaffold (Figure 2e). The EDS detected
considerable amount of calcium and phosphorus from these deposits, which
confirmed the formation of CaP (Figure 2f). These flake-like CaP depositions
formed a dense multi-layer crystal structure when the immersion period reached
4 weeks (Figure 3f). EDS also confirmed a more pronounced signal of calcium and
phosphorus on the PU/HA40 scaffolds (Figure 3e). By contrast, only limited CaP
crystal deposition was observed to scatter on the surface of PU and PU/HA20
scaffolds after 4 weeks (Figure 3a-d).
50
Biological Evaluation of PU/HA scaffolds
3
Figure 3. Biomineralization of the scaffolds after 4-week immersion in SBF. A limited
CaP crystal deposit was shown on the surfaces of PU (a and b) and PU/HA20 (c and d)
scaffolds, while a dense multi-layer crystal structure with pronounced calcium and
phosphorus composition was detected on the PU/HA40 scaffolds (e and f).
3.3 Cell viability and proliferation
As assessed by the Live/Dead assay, more than 95% of the cells were viable on
all experimental scaffolds after 24 h of cell seeding (Figure 4). Furthermore, the
DNA contents (Figure 5a), associated with the cell numbers, increased gradually
until day 8 for all groups. On day 16, a decreased DNA content was detected on all
groups compared to that on day 8, while the amount maintained stable thereafter.
At all time points, the PU/HA20 and PU/HA40 groups displayed a significantly
enhanced DNA content compared to the PU group. From day 16 onwards, a higher
DNA content was also found on group PU/HA40 compared to PU/HA20.
51
Chapter 3
Figure 4. Cell viability on PU(a), PU/HA20 (b) and PU/HA40 (c) scaffolds. More than
95% of the cells were viable on all experimental scaffolds after 24 h of cell seeding.
(Live cells are stained green and dead cells are stained red.)
Figure 5. DNA contents (a) and ALP activity (b) on PU, PU/HA20 and PU/HA40
scaffolds. *: p<0.05, **: p<0.01; error bars represent standard deviation (n=3).
52
Biological Evaluation of PU/HA scaffolds
3.4 Osteogenic differentiation
The expression of ALP activity was measured as a marker of osteogenic
differentiation of the cells (Figure 5b). Higher ALP activity was found on the PU
scaffolds compared to PU/HA composite in the early stage of cell culture until
day 8. However, the PU/HA40 scaffolds exhibited considerably higher ALP activity
compared to that on PU scaffolds at day 16 and day 24. On RNA level, the PCR
results also revealed that the PU/HA composites, compared to the PU scaffold,
promoted the expression of Bsp and Runx2 at day 4 (Figure 6a), and the expression
of Oc at day 8 (Figure 6b).
Figure 6. Expression of osteogenic related genes. Compared to PU, the PU/HA
scaffolds promoted the expression of Bsp and Runx2 at day 4 (a), and Oc at day 8 (b).
*: p<0.05, **: p<0.01; error bars represent standard deviation (n=4).
3.5 In vivo evaluation
All animals remained in good health and no signs of wound complications were
observed postoperatively. After 8 weeks, all implanted scaffolds were retrieved.
Neither macroscopic signs of inflammation nor adverse tissue responses were
discerned. Histological observation showed that all scaffolds (with or without cell
seeded) were encapsulated by a thin fibrous layer (3-6 layers of fibroblasts) without
significant inflammatory cells infiltration (Figure 7). Inside of the fibrous capsule,
new bone formation (stained red) was observed in all cell-scaffold constructs for
both PU and PU/HA40 groups, which was present not only at the periphery of the
scaffolds (Figure 7a and b), but also penetrated in the core areas (Figure 8). By
contrast, the implanted cell-free PU/HA scaffolds only exhibited fibrous capsulation
without any bone formation (Figure 7c). Due to the deformation of the scaffolds
during histological processing and the transparency of the PU scaffolds after being
53
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Chapter 3
Figure 7. Capsulation and bone formation on the periphery of the scaffolds after in
vivo implantation. All implanted scaffolds (with or without cell seeded) were
encapsulated by a thin fibrous layer without significant inflammatory infiltration.
Inside of the fibrous capsule, new bone formation (stained red) was observed in cellscaffold constructs for both PU/HA40 (a) and PU (b) groups, but not in the cell-free
PU/HA scaffolds (c). (S: scaffolds; black arrow: fibrous capsule)
embedded in MMA, quantification of the bone volume on the scaffolds could not
be performed to determine which group had higher amount of bone formation.
Nevertheless, similar bone quality was observed on the PU and PU/HA40 scaffolds
from the histological evaluation. As shown in Figure 8a and b, lamellar like bone
tissue was observed inside of the pores and mainly aligned along the pore surface.
The central area displayed a bone marrow structure, which was filled with a great
number of hematopoietic cells (stained dark blue) and adipocytes (with bubblelike morphology) (Figure 8c-d). Meanwhile, an immature woven bone like structure
was also found in some pores, which showed the onset of bony matrix deposition
with randomly arranged osteocytes. Osteoblasts were observed to align on the
periphery of the newly formed bone matrix (Figure 8d).
4. Discussion
In the current study, aliphatic PU scaffolds incorporated with different amounts of
HA particles were synthesized by in situ polymerization and simultaneous foaming
method, subsequently their biological performance for bone tissue engineering
applications was evaluated.
Biomineralization of the PU and PU/HA scaffolds, which is related to bone-bonding
capacity, was evaluated by examining the apatite formation ability on the scaffold
surfaces by incubating the scaffolds in SBF with ion concentrations equal to human
blood plasma [15]. Our results indicated that, the incorporation of 40% HA particles
significantly enhanced the biomineralization ability of the PU scaffolds. This might
be attributed to the partial dissolution of HA and the subsequent release of calcium
ions from PU/HA40 scaffold, which was more sufficient to favor CaP deposition
compared to that from the PU or PU/HA20 scaffolds. Additionally, the exposure
54
Biological Evaluation of PU/HA scaffolds
3
Figure 8. Bone in-growth to the scaffolds after in vivo implantation. On both PU/
HA40 (a and c) and PU (b and d) constructs with cell seeded, lamellar like bone tissue
(stained red) was observed in the core area of the scaffolds, mainly aligning along the
pore surface. Surrounded by the bone matrix, a large number of hematopoietic cells
(stained dark blue) and adipocytes (with bubble-like morphology) were observed,
which displayed a bone marrow structure. (B: bone; BM: bone marrow; S: scaffold)
of HA particles on the surface of PU/HA40 provided more nucleation sites for CaP
formation and growth.[16]
Biocompatibility is a primary feature of scaffold materials for tissue engineering
applications. Recently, Williams has re-evaluated the definition of biocompatibility
and pointed out that, biocompatibility of a scaffold should not be solely dependent
on not eliciting any undesirable effects on the cells or the host as an insertable
material [17]. More importantly, the scaffolds should also support the appropriate
cellular activity to optimize tissue regeneration.[17] Our results showed that both
PU and PU/HA scaffolds maintained the viability and supported progressive
growth of the seeded cells in vitro. Notably, loading HA particles into the PU
matrix improved the proliferation of the cells. The enhanced cell growth on the
PU/HA composite might be closely related to the HA particles which increased
the initial anchoring and spreading of serum proteins on the polymer surfaces.[18]
55
Chapter 3
Moreover, the addition of HA could increase surface oxygen,[19] which has been
shown to improve the attachment and the proliferation of osteoblast-like cells on
biomaterials.[20]
An enhanced osteogenic differentiation of the seeded cells has also been
demonstrated with the PU/HA scaffolds by the increase of ALP activity and the upregulation of the osteogenic related genes. The up-regulated genes were Bsp and
Runx2, which are important for orienting osteoprogenitors toward osteo-lineage
and regulating the initial stages of crystal growth,[21] and Oc, which is closely
related to the late mineralization process. Addition of HA promoting osteogenic
differentiation of the cells is probably attributed to two reasons: 1) the increased
roughness of the PU/HA scaffolds, and 2) the release of calcium and phosphate
from the PU/HA scaffolds into the cell culture medium or the microenvironment
of the seeded cells. Previous studies have shown that an increase roughness in
three-dimensional scaffolds resulted in an enhanced osteogenic differentiation.
[22]
However, such topography is expected to have a limited impact when a highly
porous and interconnected scaffold is applied, as it will be only effective for the
cells directly attached on the rough surface but not those present in the pores.[23]
Thus we speculate that the calcium and phosphate release from HA represents a
more effective differentiation signal. Our data is also consistent with other studies,
which showed that HA provides a favorable environment for osteoblast-like cell
differentiation.[24, 25]
The engineered cell-scaffold constructs were further implanted subcutaneously
to gain insight in their role to support osteogenic differentiation of the cells after
transplantation. The reason to choose an ectopic model instead of an orthotopic
model for this study was to eliminate or reduce the number of variables involved
in bone formation present in a bony environment, for instance bone stimulating
cytokines, endogenous progenitor/stem cells and potentially bone-stimulating
mechanotransduction.[26] Our result indicated that the PU and PU/HA40 scaffolds
supported the cellular growth in vivo. Furthermore, the cell-scaffold constructs
worked as an osteoinductive complex and were capable of generating new bone
tissue. After 8 weeks of implantation, the bone was still in an active forming and
remodeling phase.
On the other hand, no obvious degradation of these two scaffolds was observed
after 8 weeks of in vivo implantation. In theory, the PU matrix can undergo
degradation through the hydrolysis of ester linkages which were introduced
by castor oil.[4] The breakdown of these ester bonds yields hydroxyl and
carboxylic groups.[7] The acidic carboxyl group accelerates further hydrolysis
and the degradation becomes autocatalytic.[6] Previous results demonstrated an
56
Biological Evaluation of PU/HA scaffolds
approximate 10% weight loss of aliphatic PU/HA40 scaffolds after being soaked in
PBS for 8 weeks in vitro.[12] In our study, however, no significant degradation was
observed after 8 weeks of implantation. The possible reason is, due to the in vivo
fluid exchange and transportation, carboxyl products were easily removed from
the local environment of the scaffolds, thereby impeding the autocatalyzation
process and slowing down the progression of degradation. Considering the bone
was still in an active forming and remodeling stage, such degradation rate ensured
the PU/HA scaffolds providing sufficient mechanical support for the neotissue
before it can fulfill the structural role.[27] Still, further investigations are necessary
to evaluate the degradation of the PU/HA scaffolds in a longer term and monitor
their replacement by the tissue in-growth.
Although the PU/HA40 scaffolds showed higher biomineralization ability and
significant enhancement of osteoblastic differentiation of the seeded cells in vitro
compared to the PU scaffolds, similar quality of the ectopic bone formation was
revealed on these two scaffolds after 8 weeks in vivo. Previous studies have also
shown that there was a lack of correlation and predictability of in vitro osteogenic
marker expression on subsequent in vivo ectopic bone formation.[23, 28] One possible
reason might be that, the release of calcium and phosphate from PU/HA40 was
not sufficient to influence the cellular behaviors and local ionic concentration to
trigger more bone formation or faster bone maturation once implanted. On the
other hand, it should be noted that, the ectopic bone formation occurred in an
intradermal environment, which lacks of naturally bone-forming stem cells and
stimulators [26]. Such an environment also differs from the condition used for in vitro
cell culture where the osteogenic growth factors and supplements are sufficiently
supplied. Therefore, it is expected that the PU/HA40 scaffolds will probably
show superior behavior in promoting the osteoblastic differentiation of the
osteoprogenitors and/or stem cells when applied in the orthotopic locations. Our
study warrants further investigation towards a clinical implementation of the PU/
HA scaffolds, including 1) scaling up of the implanted constructs and 2) orthotopic
implantation in an immunocompetent animal model in which endogenous stem
cells, mechanical load and biochemical/inflammatory factors are closely involved
in the bone forming process.
5. Conclusions
Porous aliphatic PU and PU/HA composite scaffolds were synthesized by a
foaming method and their biological performances were evaluated in this study.
The incorporation of 40wt% HA nanoparticles into PU significantly promoted the
biomineralization ability of the scaffolds and enhanced the in vitro proliferation
57
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Chapter 3
and osteogenic differentiation of the seeded MSCs. In vivo implantation
revealed that a considerable amount of vascularized bone tissue with marrow
stroma development was generated on both PU and PU/HA40 scaffold after 8
weeks. Overall, with improved mechanical strength and efficacy of supporting
osteogenesis, the PU/HA composite is a potential scaffold for bone regeneration
applications.
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26. Scott, M.A., Levi, B., Askarinam, A., Nguyen, A., Rackohn, T., Ting, K., Soo, C., and James, A.W.,
Brief review of models of ectopic bone formation. Stem Cells Dev, 2012. 21(5): p. 655-67.
27. Hutmacher, D.W., Scaffolds in tissue engineering bone and cartilage. Biomaterials, 2000. 21(24):
p. 2529-43.
28. Chai, Y.C., Roberts, S.J., Desmet, E., Kerckhofs, G., van Gastel, N., Geris, L., Carmeliet, G.,
Schrooten, J., and Luyten, F.P., Mechanisms of ectopic bone formation by human osteoprogenitor
cells on CaP biomaterial carriers. Biomaterials, 2012. 33(11): p. 3127-42.
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3
Chapter 4
In vivo bone generation via
the endochondral pathway on
three-dimensional electrospun fibers
Wanxun Yang, Fang Yang, Yining Wang,
Sanne K. Both and John A. Jansen
In Vivo Endochondral Bone Generation on Wet-electrospun Fibers
Summary of this chapter
A new concept of generating bone tissue via the endochondral route might
have superiority over the standard intramembranous ossification approach.
To implement the endochondral approach, suitable scaffolds are required to
provide a three-dimensional (3D) substrate for cell population, differentiation
and eventually for the generation of osteochondral tissue. Therefore, a novel
wet-electrospinning system, using ethanol as collecting medium, was exploited in
this chapter to fabricate a cotton-like poly(lactic-co-glycolic acid) (PLGA)/poly(εcaprolactone) (PCL) scaffold that consisted of an accumulation of fibers in a very
loose and uncompressed manner. On these scaffolds rat bone marrow cells were
seeded and chondrogenically differentiated in vitro for 4 weeks followed by
subcutaneous implantation in vivo for 8 weeks. Cell pellets were used as a control.
The glycosaminoglycan (GAG) assay and Safranin O staining showed that the cells
infiltrated throughout the scaffolds and deposited an abundant cartilage matrix
after in vitro chondrogenic priming. Histological analysis of the in vivo samples
revealed extensive new bone formation through the remodeling of the cartilage
template. In conclusion, using the wet-electrospinning method, we are able to
create a 3D scaffold in which bone tissue can be formed via the endochondral
pathway. This system can be easily processed for various assays and histological
analysis. Consequently, it is more efficient than the traditional cell pellets as a tool
to study endochondral bone formation for tissue engineering purposes.
1. Introduction
Bone lesions and defects present a significant problem in clinics due to the limited
self-healing capacity of the bone tissue. Therefore, the use of a biodegradable
scaffold provided with multipotent cells and biological cues offers a promising
approach for bone repair and regeneration [1]. Since the introduction of this concept,
experiments to engineer bone tissue have been primarily focused on a process
resembling intramembranous ossification, i.e. direct osteoblastic differentiation
[2]
. However, the success of this approach is hindered by poor vascularization
inside the entire cell-based construct [3]. In line with skeleton development and
bone fracture healing, a new tactic has been reported recently, which is based
on mimicking endochondral bone formation. Thereby, the bone is generated
after a cartilage intermediate stage instead of direct osteoblastic differentiation
[4-7]
. The rationale behind this approach is that chondrocytes are able to survive
with limited nutrition and oxygen. Secondly, they can secrete vascular endothelial
growth factor (VEGF) in the hypertrophic stage, which is beneficial for blood vessel
in-growth [8]. As a consequence, this route has the potential to circumvent the
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issue faced by the conventional bone tissue engineering, that is upon implantation
the constructs will be cut off from the supply of nutrition and oxygen.
To study this endochondral pathway, the cell pellet culture system is so far the
most frequently used technique [9]. The cell pellet system provides a threedimensional (3D) environment and allows cell-cell interactions, which are similar
to pre-cartilage condensation during embryonic development. However, this
model is not suitable for the regeneration of bone tissue because of limitations
in the size, quantity, and lack of mechanical stability [10]. As a consequence, for
final clinical application of this endochondral pathway a scaffold material has to be
used, which can serve as a carrier for the cells to provide a 3D environment as well
as mechanical support. In this way, a robust cell proliferation and differentiation
can be achieved, which leads to the formation of proper extracellular matrix (ECM)
and eventually functional tissue.
Tremendous efforts have been made in the development of bone supporting
scaffolds with different compositions and 3D configurations using a wide variety
of techniques, among which electrospinning has recently gained popularity.
Electrospinning, also named electric spinning, is known as a simple and versatile
method to produce fibrous polymeric meshes [11]. With the range of tens of
nanometers to a few microns in diameter and the form of non-woven structure,
electrospun fibers are morphologically similar to natural ECM, which makes them
attractive for cells to populate on and to function effectively [12, 13]. However, a
scaffold made by the conventional electrospinning process is mostly composed
of a compact fiber network; thus it is very difficult for the cells to penetrate into
the internal part of the scaffold [14, 15]. Such a scaffold cannot be considered a 3D
structure as the cells only populate on the superficial surface.
To address the aforementioned issues, the aim of this study was to fabricate a 3D
scaffold using a modified electrospinning technique, and to further test its efficacy
in endochondral bone formation. To this end, the newly developed scaffolds were
seeded with rat bone marrow cells (RBMCs), followed by in vitro chondrogenic
differentiation and in vivo subcutaneous implantation. It was hypothesized that
the RBMCs can generate bone in vivo on the 3D electrospun scaffolds via the
endochondral pathway.
2. Materials and methods
2.1. Fabrication of electrospun scaffolds
Poly(lactic-co-glycolic acid) (PLGA, Purasorb® PDLG 8531) and poly(ε-caprolactone)
(PCL, LACTEL® Absorbable Polymers, inherent viscosity range: 1.0 – 1.3 dl/g) were
64
In Vivo Endochondral Bone Generation on Wet-electrospun Fibers
purchased from Purac Biomaterials BV (Gorinchem, The Netherlands) and DURECT
Corporation (Pelham, USA), respectively, and used in the electrospinning process.
Organic solvent 2,2,2-trifluoroethanol (TFE) (purity≥99.8%) was obtained from
Acros (Geel, Belgium). The electrospinning solution was prepared by dissolving
PLGA/PCL (weight ratio 3:1) in TFE at a concentration of 0.12 g/ml.
The 3D scaffolds were fabricated using a so-called wet-electrospinning technique
as depicted in Fig. 1 in a commercial available electrospinning set-up (Esprayer
ES-2000S, Fuence Co., Ltd, Tokyo, Japan). The optimal processing parameters for
stable formation of electrospun fibers were selected based on an earlier pilot
study. Briefly, the prepared polymer solution was fed into a syringe and delivered
to an 18G nozzle at a feeding rate of 50 μl/min. A high voltage of 20 kV was applied
at the nozzle to generate a stable polymer jet by overcoming the surface tension
of the polymer solution [16]. Different from the conventional electrospinning setup,
Figure 1. Illustration of 3D scaffold generation by the wet-electrospinning method. A
polymer solution was fed into a syringe and a high voltage was applied at the nozzle
to generate the polymer jet. A grounded bath filled with 100% ethanol was used as
collector. As ethanol is a wetting agent for both PLGA and PCL, the resulting fibers
formed a loose cotton-like mesh in the bath.
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a grounded bath filled with 100% ethanol was used to collect the fibers, located
15 cm under the nozzle. To control the size of resulting fiber meshes, the process
was stopped every 10 minutes for fiber mesh collection. Subsequently, the wetelectrospun scaffolds were washed thoroughly in MilliQ water and freeze-dried for
3 days. For comparison, electrospun fibers were also collected by the conventional
electrospinning method on top of flat aluminum foil, using the same processing
parameters and collecting time.
2.2. Characterization of the scaffolds
The morphology of the wet- and conventional electrospun scaffolds was observed
by scanning electron microscopy (SEM; JEOL6340F, Tokyo, Japan) after being
sputter-coated with gold-platinum. The fiber diameters were measured from the
SEM micrographs as obtained at random locations (n=30) using Image J software
(National Institutes of Health, Bethesda, USA).
The porosity of scaffolds was evaluated by a gravimetric measurement [17]. In brief,
the electrospun scaffolds (n = 4) were punched into disc shape forms (a diameter
of 6 mm for wet-electrospun scaffolds and a diameter of 15 mm for conventional
scaffolds) and their dimensions were measured to calculate the volume of the
scaffolds. The weight of the scaffolds was also measured to determine the
apparent density of the scaffolds (ρap). The porosity was then calculated according
to the following equation:
Porosity = (1 - ρap/ ρm) × 100%
where ρm is the density of the blend PLGA/PCL and was calculated as 1.215 g/cm3
based on the weight ratio and respective densities of PLGA (ρ = 1.24 g/cm3, Purac
MSDS database) and PCL (ρ = 1.145 g/cm3, Sigma-Aldrich MSDS database).
2.3. Cell isolation, seeding and culture
RBMCs were isolated from 7-week-old male Fischer rats after the approval from
Radboud University Nijmegen Animal Ethics Committee (Approval no: RU-DEC
2011-142). Briefly, two femora of each rat were removed and the epiphyses
were cut off. RBMCs were flushed out of the remaining diaphyses using the
proliferation medium through an 18G needle (BD Microlance, Drogheda, Ireland).
The proliferation medium consisted of alpha Minimal Essential Medium (aMEM;
Gibco, Grand Island, USA) supplemented with 10% fetal bovine serum (FBS;
Sigma, St. Louis, USA), 100 U/ml penicillin, 100μg/ml streptomycin (Gibco), 50
μg/ml l-ascorbic acid (Sigma) and 10nM dexamethasone (Sigma). The flush-out
was cultured for 2 days in a humidified incubator (37 °C, 5% CO2), after which the
medium was refreshed to remove non-adherent cells. Prior to detaching the cells,
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In Vivo Endochondral Bone Generation on Wet-electrospun Fibers
they were cultured for an additional 3 days then the cells were detached using
trypsin/EDTA (0.25% w/v trypsin, 0.02% w/v EDTA; Sigma) and counted.
Disk-shaped scaffolds with a diameter of 6 mm and a thickness of about 2.5
mm were punched out (biopsy punch; Kai medical, Gifu, Japan) from each wetelectrospun mesh that was collected for 10 minutes. Afterwards they were
sterilized in 70% ethanol for 2 hours and soaked in the proliferation medium
overnight after washed with PBS. During cell loading, the scaffolds were incubated
in the cell suspension with a concentration of 1 × 106 cells/ml (4 scaffolds per 1 ml
of cell suspension) and gently rotated for 3 hours. Successively, the scaffolds were
placed in non-adherent tissue culture plates and the unattached cells from the
suspension were collected, centrifuged and reseeded onto the scaffolds to ensure
a high cell loading efficiency. The cell-scaffold constructs were then cultured
for 1 week in proliferation medium. Thereafter, the medium was changed to
chondrogenic medium, consisting of high-glucose DMEM (Gibco), 1% FBS, 100 U/
ml penicillin, 100 μg/ml streptomycin, 50 μg/ml l-ascorbic acid, 100 mM of sodium
pyruvate (Gibco), 1:100 insulin-transferrin-selenium (ITS; BD Biosciences, Bedford,
USA), 100nM dexamethasone, 10 ng/ml of transforming growth factor beta-2
(TGF-β2; R&D systems, Minneapolis, USA) and 100 ng/ml bone morphogenetic
protein 2 (BMP-2; BD biosciences). The constructs were cultured for 4 weeks
and the medium was refreshed twice a week. As a control, also a pellet culture
was made to induce chondrogenic differentiation. To ensure a similar amount
of cell doublings compared to the constructs, the cells were re-seeded after the
first passage in tissue culture flasks and kept in proliferation medium for 1 week.
Subsequently, cell pellets (250,000 cells/pellet) were made by centrifugation for 8
minutes at 200g and cultured in the chondrogenic medium for 4 weeks. Both the
cell-scaffold constructs and the pellets were then used for in vitro and in vivo tests.
2.4. In vitro studies
After 4 weeks of chondrogenic differentiation, both the cell-scaffold constructs and
the cell pellets were tested for their glycosaminoglycan (GAG) and DNA content. A
Safranin O staining was performed to analyze the in vitro formed cartilage.
The amount of GAG in each sample was detected by a sulfated glycosaminoglycan
(sGAG) assay (n=3). After being washed twice in PBS, all samples were digested
in proteinase K solution (1 mg/ml proteinase K in 50 mM Tris with 1 mM EDTA, 1
mM iodoacetamide, and 10 μg/ml pepstatin A) (all reagents from Sigma) for 16
hours at 56°C. The GAG content was determined by using dimethylmethylene blue
(Polysciences Inc., Warrington, USA), with chondroitin sulfate (Sigma) as standard
control. All results were measured with an ELISA reader (Bio-Tek Instruments Inc.,
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Winooski, USA) at 530 nm and 590 nm. For each sample, the GAG content was
normalized by the amount of DNA as described below.
DNA content (n=3) was quantified by PicoGreen assay (Quant-iT PicoGreen dsDNA,
Invitrogen, Oregon, USA) using the same cell extract solution as used for the GAG
assay. A DNA standard curve was made and the results were obtained using a
fluorescence microplate reader (Bio-Tek Instruments Inc.) with an excitation
wavelength at 485nm and an emission wavelength at 530nm.
For the Safranin O staining, the in vitro scaffolds (n=4) were fixed in 10% phosphate
buffered formalin for 1 hour and embedded in paraffin. The microtome sections
with 6 μm of thickness were stained with Safranin O (Merck, Darmstadt, Germany)
after deparaffinization in xylene and rehydration through graded ethanol.
2.5. In vivo implantation
After 4 weeks of in vitro culture, the cell-scaffold constructs and the cell pellets
were implanted subcutaneously in ten 8-week-old male nude rats (Crl:NIH-Foxn1mu,
Charles River). The protocol was approved by the Animal Ethical Committee of the
Radboud University Nijmegen Medical Centre (Approval no: RU-DEC 2011-142) and
the national guidelines for the care and use of laboratory animals were applied.
Surgery was performed under general inhalation anesthesia with a combination
of isoflurane, nitrous oxide, and oxygen. The back of the rats was shaved and
disinfected with povidone-iodine. Subsequently, two small longitudinal incisions
were made. Lateral to each incision, a subcutaneous pocket was created. After
placing the samples, the skin was closed using staples. One cell-scaffold construct
or three cell pellets were inserted in each pocket. After 8 weeks, the rats were
euthanized by CO2-suffocation for sample collection.
2.6. Histological staining
All in vivo samples were fixed in 10% phosphate buffered formalin for 36 hours and
dehydrated through graded ethanol before embedding in either methylmetacrylate
(MMA, L.T.I., Bilthoven, the Netherlands) or paraffin.
MMA embedded samples (n = 5) were sectioned using a sawing microtome
technique (Leica SP1600, Leica microsystem, Wetzlar, Germany) and stained with
methylene blue and basic fuchsin (both reagents from Merck).
Prior to paraffin embedding, the fixed samples (n = 5) were decalcified by 10%
EDTA and the decalcification process was monitored via X-ray. Six μm-thick sections
were made from each sample and used for hematoxylin/eosin (HE), elastic Van
Gieson (EVG) and Safranin O staining.
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In Vivo Endochondral Bone Generation on Wet-electrospun Fibers
2.7. Back-scattering scanning electron microscopy (BS-SEM)
Thick MMA sections (n=3) were polished by silicon carbide sandpapers up to grit
4000 and then sputter-coated with gold–platinum. The micrographs were acquired
by high resolution SEM in the backscattering mode at 25kV (JEOL6340F, Tokyo,
Japan).
2.8. Statistic analysis
All data in this study was reported as means ± SD. Statistical significance was
evaluated using an unpaired t-test (Instat 3.05, GraphPad Software Inc., San Diego,
USA). Differences were considered statistically significant at p < 0.05.
3. Results
4
3.1. Scaffold characterization
The wet-electrospinning technique developed in this study displayed obvious
differences in the collected fiber structures compared with conventional
electrospinning. After 10 minutes of collection, the wet-electrospun PLGA/PCL
mesh showed a loose and fluffy structure with a diameter of 3-4 cm and a thickness
of 2-3 mm (Fig. 2a, I). In contrast, the fibers collected via the conventional way
formed a thin and compressed membrane with a diameter of around 12 cm and
a thickness of 40 μm (Fig. 2a, II). The SEM micrographs show that the fibers were
oriented in a random and dispersive manner, forming a porous structure for both
types of scaffolds (Fig. 2b and 2c). The average fiber diameters measured from the
SEM micrographs are 1.98 ± 0.51 µm for the wet-electrospun scaffolds and 2.01 ±
0.56 µm for the conventional electrospun scaffolds, with no statistical significant
difference. However, there exists a significant difference between the measured
porosities, showing 99.0 ± 0.2% for the wet-electrospun scaffolds and 79.4 ± 2.8%
for the conventional scaffolds, respectively.
3.2. DNA and GAG content
RBMCs were seeded onto the scaffolds to test their chondrogenic potential
compared to the cell pellet system. Prior to in vivo implantation the amount of
DNA and deposited GAG in the pellets and scaffolds were measured. As shown
in Fig. 3a, a significantly higher amount of DNA could be found in the PLGA/PCL
scaffolds compared to the cell pellets after 4 weeks of chondrogenic differentiation
in vitro. Moreover, the amount of GAG formed on the PLGA/PCL scaffolds, after
normalization by the DNA amount, was five times higher compared to that of the
cell pellets (Fig. 3b).
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Chapter 4
Figure 2. Comparison of different PLGA/PCL scaffolds. The cotton-like scaffold made
via the wet-electrospinning method (a, I) showed an uncompressed structure and
increased thickness compared to the membrane made via the conventional method
(a, II). Dispersive and randomly oriented fiber arrangement was observed by SEM for
both wet-electrospun scaffolds (b) and conventional electrospun scaffolds (c),
showing a porous structure.
3.3. Histological analysis of in vitro cartilage formation
To visualize the cartilage matrix deposition in the cell pellets and the scaffolds,
sections were made and stained with Safranin O after 4 weeks of chondrogenic
culture. As shown in Fig. 3e, the cell-scaffold constructs displayed an abundant
GAG content stained strongly in orange-red color, representing the cartilage
matrix. The PLGA/PCL fibers were embedded within the cartilage matrix and a
homogeneous distribution of cells could be observed inside of the fibrous structure.
The cell morphology differed due to the different stages of cell differentiation. A
proportion of cells displayed fibroblast-like morphology, while a great number of
cells had a typical chondrocyte morphology embedded in large lacunae and formed
isogenous groups (Fig. 3f). In contrast, the cell pellets only displayed chondrogenic
differentiation in the outer layer. There was no cartilage matrix found in the central
region and a number of cells displayed a necrotic phenotype, leaving empty spaces
inside of the pellets (Fig. 3c and d).
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In Vivo Endochondral Bone Generation on Wet-electrospun Fibers
3.4. Histological analysis of in vivo bone formation
After 8 weeks of implantation, an increased hardness of the scaffolds could be
sensed during sample retrieval. Unfortunately, the implanted cell pellets could not
be traced. To evaluate the bone forming quality and quantity, different histological
staining methods were used for either MMA or paraffin embedded samples.
4
Figure 3. DNA content and chondrogenic differentiation. After 28 days of chondrogenic
differentiation in vitro, a significantly higher amount of DNA (a) and GAG (b,
normalized for DNA), was found in the PLGA/PCL scaffolds compared to the cell
pellets. In the cell pellets, only a limited region in outer layer underwent chondrogenic
differentiation (c and d). In PLGA/PCL scaffolds, abundant GAG (orange-red color)
was secreted, which was visualized by a Safranin O staining (e). Cells with different
morphologies could be observed; chondrocyte-like cells were embedded in lacunae
and formed isogenous groups (f, black arrows), while a proportion of cells kept their
original fibroblast-like morphology (f, white arrows). ** p<0.01; error bars represent
standard error of the mean (n=3).
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Chapter 4
Four out of five MMA embedded scaffolds showed bone formation. The bone
forming quality and quantity were similar in the four samples. As shown in Fig. 4a,
extensive bone formation can be observed inside the scaffolds. Islets of bone tissue
are scattered amongst the remnants of scaffold material, indicating the onset of
the newly formed bone (Fig. 4b). Concentrically organized bone structures which
indicate mature bone could hardly be seen in the bone forming area.
Figure 4. In vivo bone formation in PLGA/PCL electrospun scaffolds. Sections of
samples after 8 weeks in vivo were stained with methylene blue and basic fuchsin.
Bone formation (red) occurred inside of the scaffold (a). The osteocytes embedded in
the bone matrix showed a random arrangement (b, white arrows). The onset of the
newly formed bone tissue (b, black arrows) was scattered amongst the remnants of
scaffold material.
To better discriminate tissue structures within the scaffolds, HE staining was applied
to the paraffin embedded samples. All five scaffolds demonstrated bone formation
with similar quality and quantity. As shown in Fig. 5a, woven bone like structure
can be observed throughout the scaffold. A large amount of immature osteocytes
with round shapes were embedded and arranged randomly, while chondrocytes
were hardly seen at this time point. The newly formed bone resembled a mixture
of cartilage and calcified cartilage tissues (bone). The vascular invasion could be
observed among the material remnants and the fibrous tissue (Fig. 5d).
To further investigate the state of bone maturation, an EVG staining was performed.
Thereby the bone area and the surrounding fibrous tissue, as well as the different
maturities of the new bone, were distinguished by the clear contrast between
the red, pink and light brown colors, representing bone, uncalcified cartilage and
fibrous tissue, respectively (Fig. 5c).
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In Vivo Endochondral Bone Generation on Wet-electrospun Fibers
4
Figure 5. Histological stainings of PLGA/PCL electrospun scaffolds after in vivo
implantation. After 8 weeks of implantation, HE staining (a and d) showed an
extensive distribution of woven bone like structure and blood vessel infiltration (d,
black arrows). Safranin O staining (b and e) confirmed the existence of remaining
GAG (red color) (e, black arrow) among calcified cartilage (green color) (e, asterisk).
Bone (f, asterisk), uncalcified cartilage (f, black arrow) and the surrounding fibrous
tissue were distinguished by the contrast between red, pink and light brown color in
EVG staining (c and f).
Furthermore, a Safranin O staining (Fig. 5b and e) confirmed the existence of the
remaining GAG (red color) among calcified cartilage (green color). Compared
to the cell-scaffold constructs before implantation (Fig. 3e), the in vivo samples
consisted extracellular matrix with considerably reduced GAG levels. Although a
few sporadic cells with strong red color remained, most regions of the cartilage
matrix were remodeled and calcified into bone tissue.
3.5. BS-SEM analysis
The grey levels in the BS-SEM images showed that there is a clear difference
between the electrospun fibers and newly formed bone (Fig. 6a). Bone tissue, with
the highest density, displayed by a grey opaque in the black background of the
MMA block. A great number of electrospun fibers were distributed dispersively
inside the bone area, visualized by black round dots or needle-like shapes,
depending on the sectioning angles (Fig. 6b). Moreover, the differences in bone
density could also be addressed according to the grey scales. Immature bone was
found in most of the bone forming areas displayed by a relatively darker color.
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Chapter 4
4. Discussion
In this study, we successfully developed a simple wet-electrospinning system by
using ethanol as the collecting media to fabricate a highly porous 3D electrospun
scaffold for bone tissue engineering applications. Successively, the capacity of
RBMCs to generate de novo bone tissue in vivo through the endochondral pathway
was investigated using this 3D scaffold. After 4 weeks of in vitro chondrogenic
priming, a cell infiltration throughout the scaffolds and abundant cartilage matrix
formation were observed. Upon subcutaneous implantation, the bone tissue was
generated ectopically by extensive calcification of the cartilage templates with
significant blood vessel infiltration.
Figure 6. Back scattering SEM images of PLGA/PCL electrospun scaffolds after in vivo
implantation. In the black background of the MMA (M), bone forming area (B)
demonstrated grey opaque with a great number of electrospun fibers (black color)
penetrating. In the magnified image of the red squared box (b), bone tissue with
different densities showed in different grey scales; brighter colors indicated the
regions of more mature bone.
Essentially, the lack of a real 3D structure in conventional electrospun scaffolds
hampers their use in tissue regeneration. To overcome this drawback, various
methods have been recently explored. A salt leaching/gas forming technique was
used to introduce micro-voids into the scaffold after electrospinning [18]. However,
the fiber morphology of the scaffolds could not be well preserved after the postspinning processes. Ultrasonication of the electrospun membrane in an aqueous
solution was also applied to make the densely packed fibers become loose, while
the increase in pore size and the thickness of the membrane was limited [19].
Pham et al. designed bilayered constructs consisting of microfibers with varying
thicknesses of nanofibers on top to improve the cell infiltration [17]. However,
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In Vivo Endochondral Bone Generation on Wet-electrospun Fibers
this method implied an increased complexity of the technique. To modify the
collecting apparatus, Blakeney et al. embedded a metal array in a semi-spherical
dish which allowed the fibers to accumulate in an open space and formed a
focused and uncompressed 3D scaffold, resembling a cotton ball [20]. It was shown
that such a scaffold allowed better cell penetration and proliferation compared
to the conventional electrospun scaffolds. In this study, a similar cotton ball like
structure was made by wet-electrospinning technique, simply using ethanol as the
collecting media. The general technique of electrospinning into a coagulation bath
was first introduced by Yokoyama et al. in 2009 [21]. Using ethanol as collecting
media was first reported by Fang et al. and used for studying the evolution of fiber
morphology [22]. As ethanol is a wet agent for synthetic polymers, e.g. PLGA and
PCL, it prevents the electrospun fibers from densely packing on top of each other
in contrary to the conventional flat collector. Afterwards, ethanol was replaced by
water and the sample was freeze-dried. During this process, the loose structure
could be well preserved and characterized. Moreover, the use of ethanol did
not alter the morphology and size of the electrospun fiber. To further prove the
efficacy of wet-electrospinning, fabricated 3D PLGA/PCL scaffolds were combined
with RBMCs to test their bone forming ability in vivo.
Loosened electrospun scaffolds have been used before for bone tissue engineering
purposes [23, 24]. However, these were implanted in rodent calvaria using different
methods for cell differentiation. Our study is the first instance of endochondral
bone formation subcutaneously using an electrospun scaffold as a tool to improve
efficacy. To achieve in vivo bone generation, we chose the pathway that resembles
endochondral ossification by first forming a cartilage template in vitro and
subsequently allowing the cartilage to be replaced by bone in vivo. This process
relies on the specialized functions of hypertrophic chondrocytes, from which a
type X collagen-rich avascular cartilaginous matrix can be produced in vitro.
After in vivo implantation, it is expected that the hypertrophic chondrocytes at
the periphery of the cartilage template instruct the surrounding mesenchymal
cells to differentiate into osteoblasts. In parallel, the chondrocytes in the central
regions lead mineralization of the hypertrophic cartilage by initiating remodeling
via the production of specific matrix metalloproteinases (MMPs) and attract blood
vessels by releasing VEGF [25]. Using this approach, a high number of well-organized
blood vessels were present within the scaffolds, which showed the feasibility and
superiority of endochondral route for engineering a vascularized bone construct.
As demonstrated by Chan at al [26], endochondral ossification is required for the
haematopoietic stem cell niche formation with a subpopulation of fetal progenitor
cells giving rise to bone with a marrow cavity. This happens only if they would
undergo endochondral ossification as opposed to intramembranous ossification.
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In this study, extensive and robust bone formation was demonstrated. It is known
that osteogenesis and angiogenesis are two closely correlated processes during
bone growth, remodeling and repair [27]. In the endochondral bone formation, the
released VEGF regulates blood vessel invasion into the hypertrophic cartilage [28].
Afterwards, undifferentiated mesenchymal cells are brought to the mineralization
front through the invading blood vessels thereby participating in the osteogenesis
[29]
. Thus the turnover from cartilage to bone is synergized when a good
vascularization is established inside the cell-scaffold constructs.
In this study, cell pellets were used as control because they provide a 3D
environment and high cell density for cell-cell interactions which mimic the precartilage condensation during embryonic development [30, 31]. This system has
been applied as a study model for engineering chondrogenically differentiated
constructs in vitro and also for the creation of vascularized bone in vivo [32]. The
in vitro results in our study showed greater cartilage production in PLGA/PCL
scaffolds, while the cell pellets exhibited limited chondrogenic differentiation in
the outer layer and necrosis in the core area. This could be related to the fact
that the scaffolds provide a solid surface for cell attachment while allowing better
penetration of fluid flows carrying the oxygen and nutrient for the developing
matrix in vitro. Moreover, the form of cell-scaffold system might not only mimic
the condensation phase directly but also provide some physical stimulus and
biomechanical cues related to the cellular environments in vivo [33]. During the in
vivo experiment, the implanted cell pellets could unfortunately not be retrieved.
In a previous study the pellets were retrievable when a nude mouse model was
used [4]. This time we used nude rats and the small pellets might have fused with
the thick subcutaneous adipose tissue thereby hampering sample retrieval. It is
known that the pellets occasionally get lost in an in vivo model due to their small
size. Obviously, the scaffold system provided better handling abilities. Besides the
aforementioned aspects, various assays and histological processing could also be
implemented using this PLGA/PCL scaffold, which makes it a superior model for
studying endochondral bone regeneration.
It is known that the progression of bone formation in vivo varies when the
cartilaginous templates are generated in different developmental stages in
vitro before implantation [30]. However, no conclusion has been reached on the
length of in vitro chondrogenic priming time because different cell resources or
scaffold types were applied in previous studies [4, 34-36]. In this study 4-weeks of
in vitro chondrogenic differentiation were applied. Histological staining showed
extensively cartilage matrix deposition within the scaffolds and a number of
chondrocytes presented hypertrophic morphology. In vivo results demonstrated
that 8 weeks of implantation was not sufficient to ossify the complete construct,
76
In Vivo Endochondral Bone Generation on Wet-electrospun Fibers
while a comprehensive remodeling of the cartilage has been revealed. The
new bone (woven bone) mainly consisted of immature characteristic of bone
and uncalcified cartilage. A prolonged in vivo implantation time may result in
the complete maturation and eventually the functional bone tissue formation
with marrow stroma. On the other hand, more research work is still needed to
determine the optimal in vitro priming time to obtain satisfactory bone formation
in vivo.
From material aspects, it should also be noted that PLGA and PCL, as synthetic
polymers, are in general not osteoconductive. In most cases, polymers are combined
or coated with osteoconductive components, for instance calcium phosphates
and hydroxyapatite (HA), to improve their osteo-compatibility [37]. Previous studies
verified that a composite electrospun scaffold based on HA and PCL could stimulate
bone formation by promoting cell proliferation and osteoblastic differentiation
[38]
. Therefore, the addition of osteoconductive components may offer some
extra biological cues to influence the bone forming process during endochondral
route. Additionally, the modulation of material properties of PLGA/PCL, such as
degradation time [39], may also influence the in vivo outcome. It has been found
that the incorporation of nano sized HA could slow down the degradation of PLGA/
PCL materials in a HA amount-dependent manner and significantly improved the
tissue response during 4-week of subcutaneous implantation [39]. For that reason,
it is necessary to address the relationship between the supporting material and
endochondral bone formation in future research, so that this PLGA/PCL constructs
can be eventually applied for maxillofacial or dental applications, e.g. bone
reconstruction for dental implantation and bone defects in non-load bearing sites.
5. Conclusions
In this study the new concept of engineering bone tissue via the endochondral route
was successfully combined with a 3D PLGA/PCL blend scaffold fabricated using a
wet-electrospinning system. After 4 weeks of in vitro chondrogenic priming, a cell
infiltration and abundant cartilage matrix deposition by RBMCs could be observed
throughout the scaffolds. Eight weeks of in vivo implantation revealed extensive
new bone formation through the remodeling of cartilage template. This system
can be easily processed for various assays and histological analysis. Therefore, it is
more efficient than the cell pellets as a tool to study endochondral bone formation
for tissue engineering purposes.
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4
Chapter 4
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79
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Chapter 5
Performance of different
three-dimensional scaffolds for in
vivo endochondral bone generation
Wanxun Yang, Sanne K. Both , Gerjo J. V. M.van Osch,
Yining Wang, John A. Jansen and Fang Yang
Performance of Different 3D Scaffolds for Endochondral Bone Formation
Summary of this chapter
In the context of skeletal tissue development and repair, endochondral ossification
has inspired a new approach to regenerate bone tissue in vivo via a cartilage
intermediate as an osteoinductive template. The aim of this chapter was to
investigate the behavior of mesenchymal stem cells (MSCs) in regard to in vitro
cartilage formation and in vivo bone regeneration when combined with different
three-dimensional (3D) scaffold materials, i.e. hydroxyapatite/tricalcium phosphate
(HA/TCP) composite block, polyurethane (PU) foam, poly(lactic-co-glycolic acid)/
poly(ε-caprolactone) electrospun fibers (PLGA/PCL) and collagen I gel. To this end,
rat MSCs were seeded on these scaffolds and chondrogenically differentiated in
vitro for 4 weeks followed by in vivo subcutaneous implantation for 8 weeks. After
in vitro chondrogenic priming, comparable cell amounts and cartilage formation
were observed in all types of scaffolds. Nonetheless, the quality and maturity of
in vivo ectopic bone formation appeared to be scaffold/material-dependent. Eight
weeks of implantation was not sufficient to ossify the entire PLGA/PCL constructs,
albeit a comprehensive remodeling of the cartilage had occurred. For HA/TCP,
PU and collagen I scaffolds, more mature bone formation with rich vascularity
and marrow stroma development could be observed. These data suggest that
chondrogenic priming of MSCs in the presence of different scaffold materials allows
the establishment of reliable templates for generating functional endochondral
bone tissue in vivo without using osteoinductive growth factors. The morphology
and maturity of bone formation can be dictated by the scaffold properties.
1. Introduction
Since the introduction of tissue engineering concept, cell-based bone regeneration
strategies have received much attention [1, 2]. The conventional approach in this
field is to combine mesenchymal stem cells (MSCs) with biomaterials/scaffolds
followed by directing MSCs towards the bone lineage in vitro, which resembles
intramembranous ossification [3]. However, the clinical translation of this approach
so far has been unsatisfactory [2]. Lack of sufficient vascular supply, resulting in
immediate cell death after implantation, is generally thought to be the cause of
failure of the cell-scaffold constructs in patients [4].
Recently, in the context of skeletal tissue development and repair, endochondral
ossification has inspired a new approach to regenerate bones [5-7]. In this approach,
MSCs seeded on scaffolds are first primed towards the chondrogenic lineage in
vitro before implantation. These cartilage intermediates work as osteoinductive
templates and transform into bone tissue in vivo [4, 8, 9]. The recent findings suggest
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that this endochondral ossification approach may be more successful compared
to the intramembranous route. It is known that cartilage is well adapted to limited
nutrition supply and hypoxic conditions [10]. In addition, during the hypertrophic
stage of the endochondral process, the chondrocytes produce substantial
amount of vascular endothelial growth factor (VEGF), which plays a critical role
in the establishment of vessel in-growth and extracellular matrix remodeling [11].
The morphogenetic signals from the chondrocytes also promote the osteogenic
differentiation of local MSCs [12, 13]. Hence, the endochondral bone tissue
engineering route may allow for a longer temporal window for vascularization and
has potential for engineering larger bone constructs .
For the final clinical application of such a strategy, an appropriate scaffold is an
indispensable element, which provides a three-dimensional (3D) environment
as well as the mechanical support for the cells [14]. Regarding endochondral bone
formation, it is a unique transition process from chondrogenesis to osteogenesis
and involves several instances of extracellular matrix (ECM) degradation and
remodeling. Thus, the scaffolds for endochondral bone tissue engineering should
be able to support cells in differentiating into chondrogenic lineage in vitro and
meanwhile, also promote the calcification of a cartilage template into bone after
in vivo implantation. Cartilage matrix consists of highly hydrated proteoglycan
embedded into a type II collagen network. Based on the previous findings,
hydrogels made from natural or synthetic polymers have shown to support the
phenotype of chondrocyte [15]. On the other hand, bone tissue is more rigid than
cartilage and impregnated mainly with hydroxyapatite (HA) and type I collagen.
Among the candidate scaffold materials, porous bioceramics and degradable
polymers are widely studied because of their mechanical competence and
structural feasibility for fluid perfusion and vessel in-growth. Apart from the
aforementioned, nanofibrous scaffolds made by electrospinning has also gained
research interests in both cartilage and bone tissue engineering due to their
morphological resemblance to the natural extracellular matrix [16, 17].
In this study, we aimed at testing the performance of different 3D scaffolds for
endochondral bone tissue engineering. To that end, four representative scaffolds
(ceramic- or polymer-based) with distinct structural features were selected
from the aforementioned materials, including porous hydroxyapatite/tricalcium
phosphate (HA/TCP), porous polyurethane (PU), fibrous poly(lactic-co-glycolic
acid) (PLGA) electrospun scaffolds and collagen I hydrogel. Rat primary MSCs were
seeded on these scaffolds and chondrogenically primed for 4 weeks in vitro. The
formation of cartilage templates was characterized by matrix quantification and
gene analysis. After 4 weeks of in vitro culture, the chondrogenically primed cellscaffold constructs were subcutaneously implanted in rats for 8 weeks and then
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Performance of Different 3D Scaffolds for Endochondral Bone Formation
the spatial pattern of bone formation and cartilage remains were assessed in each
construct in terms of quality and quantity via immunohistological stainings and
polarized microscopy.
2. Materials and Methods
2.1 Scaffold preparation
Four different types of scaffold materials were used in this study:
Porous sintered HA/TCP composite (CAM bioceramics B.V., Leiden, the Nethelands)
was obtained in a cubic shape of 3×3×3 mm3 and consisted of HA and TCP with
a 75/25 ratio. The volumetric porosity was 75% (provided by the manufacturer).
HA/TCP (Fig. 1a) showed a macro-porous structure with pore sizes ranging from
300-800 mm.
Porous PU scaffold was fabricated by in situ polymerization and simultaneous
foaming as previously described [18, 19]. The scaffold displayed a similar porous
structure as HA/TCP (Fig. 1b) and the porosity is 80% [18].
Figure 1. SEM micrographs of HA/TCP (a), PU (b) and PLGA/PCL (c) scaffolds. (Scale
bar in a and b: 100µm; c: 10µm) (Micrograph of collagen I scaffold is not shown
because the structure could not be preserved after freeze-drying.)
Fibrous PLGA (Purasorb® PDLG 8531, Purac Biomaterials BV, Gorinchem, the
Netherlands) combined with poly(ε-caprolactone) (PCL, LACTEL® Absorbable
Polymers, DURECT Corporation , Pelham, USA) electrospun scaffold was made
by a wet-electrospinning technique (Esprayer ES-2000S, Fuence Co., Ltd, Tokyo,
Japan) as previously described [20]. The obtained scaffold (Fig. 1c) displayed an
uncompressed structure with an average fiber diameter of 1.98 ± 0.51mm and a
porosity of 99% [20].
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Collagen hydrogel was prepared from type I collagen (rat tail) (BD Biosciences,
Bedford, USA) with the final concentration at 6 mg/ml according to the
manufacturer’s instructions.
Both PLGA/PCL and PU scaffolds were punched into a disk shape for the following
experiments with a diameter of 6 mm and a thickness of 2 mm. Collagen I gel was
shaped to the same size during the cell seeding procedure. The morphology of the
scaffolds was observed by scanning electron microscopy (SEM; JEOL6340F, Tokyo,
Japan) after being sputter-coated with gold-platinum.
2.2 Cell isolation, seeding and culture
MSCs were isolated from 6-week-old male Fischer rats (12 rats in total) (Charles
River) after the approval from Radboud University Nijmegen Animal Ethics
Committee (Approval no: RU-DEC 2011-142). Briefly, two femora of each rat
were removed and the epiphyses were cut off. Rat MSCs were flushed out of the
remaining diaphyses using the proliferation medium through an 18G needle (BD
Microlance, Drogheda, Ireland). The proliferation medium consisted of alpha
Minimal Essential Medium (αMEM; Gibco, Grand Island, USA) supplemented with
10% fetal bovine serum (FBS; Sigma, St. Louis, USA), 100 U/ml penicillin, 100μg/ml
streptomycin (Gibco), 50 μg/ml l-ascorbic acid (Sigma) and 10nM dexamethasone
(Sigma). The flush-out was cultured for 2 days in a humidified incubator (37 °C, 5%
CO2), after which the medium was refreshed to remove non-adherent cells. The
cells were cultured for an additional 3 days followed by detachment using trypsin/
EDTA (0.25% w/v trypsin, 0.02% w/v EDTA; Sigma) and counted.
Before cell seeding, the scaffolds were sterilized either by autoclave (PU and
HA/TCP) or incubation in 70% ethanol for 2 hours (PLGA/PCL) then washed
with PBS. Subsequently, all scaffolds were soaked in the proliferation medium
overnight. During cell loading, the scaffolds were incubated in the cell suspension
with a concentration of 1 × 106 cells/ml (4 scaffolds per 1 ml of cell suspension)
and gently rotated for 3 hours. Successively, the scaffolds were placed in nonadherent tissue culture plates and the unattached cells from the suspension
were collected, centrifuged and reseeded onto the scaffolds to ensure a high cell
loading efficiency. For collagen I scaffolds, the gel was made aseptically according
to the manufacturer’s instructions, mixed with cells, and set in cell culture inserts
(ThinCertTM for 24-well plate, Greiner Bio-One, Germany). A concentration of
250,000 cells per insert (with 120µl gel) was used to obtain a disc shaped cell/
collagen gel construct with a diameter of 6 mm and a thickness of 2 mm.
All cell-scaffold constructs were cultured for 1 week in the proliferation medium.
Thereafter, the medium was changed to the chondrogenic medium, consisting of
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Performance of Different 3D Scaffolds for Endochondral Bone Formation
high-glucose DMEM (Gibco), 1% FBS, 100 U/ml penicillin, 100 μg/ml streptomycin,
50 μg/ml l-ascorbic acid, 100 mM of sodium pyruvate (Gibco), 40μg/ml l-proline
(Sigma), 1:100 insulin-transferrin-selenium (ITS; BD Biosciences, Bedford, USA),
100nM dexamethasone, 10 ng/ml of transforming growth factor beta-2 (TGF-β2;
R&D systems, Minneapolis, USA) and 100 ng/ml bone morphogenetic protein
2 (BMP-2; BD biosciences). The constructs were cultured for 4 weeks and the
medium was refreshed twice a week. Afterwards the samples were divided into
two groups, one for in vitro characterization and the other for in vivo implantation.
2.3 In vitro tests
2.3.1 GAG content
The amount of GAG in each sample was detected by a sulfated glycosaminoglycan
(sGAG) assay (n=3). After being washed twice in PBS, all samples were digested
in proteinase K solution (1 mg/ml proteinase K in 50 mM Tris with 1 mM EDTA, 1
mM iodoacetamide, and 10 μg/ml pepstatin A) (all reagents from Sigma) for 20
hours at 56°C. The GAG content was determined by using dimethylmethylene blue
(Polysciences Inc., Warrington, USA), with chondroitin sulfate (Sigma) as standard
control. Measurements of absorption were performed at 530 nm and 570 nm
using an ELISA reader (Bio-Tek Instruments Inc., Winooski, USA). The results were
calculated according to the previously described method [21]. For each sample, the
GAG content was normalized by the amount of DNA as described in 2.3.2.
2.3.2 DNA content
To evaluate the seeding efficiency on each type of scaffolds and the cell number
before the implantation, DNA content of the seeded cells was measured after
24 hours of seeding and 28 days of chondrogenic differentiation, respectively.
PicoGreen assay (Quant-iTTM PicoGreen® dsDNA kit, Invitrogen, Oregon, USA)
was performed using the same cell extract solution as treated for the GAG assay.
A DNA standard curve was used to quantify the amount of DNA in each sample
and the results were measured using a fluorescence microplate reader (BioTek Instruments Inc.) with an excitation wavelength at 485nm and an emission
wavelength at 530nm.
2.3.3 RNA extraction and qPCR analysis
Total RNA was extracted using Trizol method after 4 weeks of chondrogenic culture
(n=4). Briefly, scaffolds with cells were washed with PBS and cut into small pieces
before adding 1 ml of Trizol (Invitrogen). The cell extract was collected, mixed
with chloroform and centrifuged. Only the upper aqueous phase was collected
and mixed with equal amount of isopropanol. After 10 min of incubation at room
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Chapter 5
temperature, the mixture was centrifuged and washed twice with 75% alcohol.
Thereafter, the obtained RNA pellet was dissolved in RNase free water and the
RNA concentration was measured with spectrophotometer (NanoDrop 2000,
Thermal scientific, Wilmington, USA).
Table 1. Primer sequences used for qPCR
Forward (5’ → 3’)
Reverse (5’ → 3’)
Col2α1
GACGCCACGCTCAAGTCGCT
CGCTGGGTTGGGGTAGACGC
Col10α1
GGGCCCTATTGGACCACCAGGTA
CCGGCATGCCTGTTACCCCC
Acan
CATTCGCACGGGAGCAGCCA
TGGGGTCCGTGGGCTCACAA
Vegfa
ACTCATCAGCCAGGGAGTCT
GGGAGTGAAGGAGCAACCTC
Gapdh
CGATGCTGGCGCTGAGTAC
CGTTCAGCTCAGGGATGACC
First strand cDNA was reverse transcribed from RNA using the SuperScript® FirstStrand Synthesis System kit (Invitrogen). Afterwards cDNA was further amplified
and the expression of specific genes was quantified using IQ SYBR Green Supermix
PCR kit (BioRad, Hemel Hempstead, United Kingdom) in a real-time PCR (BioRaD,
CFX96™ real-time system) (n=4). Chondrogenic differentiation related gene
markers were evaluated, including collagen type II (Col2α1), collagen type X
(Col10α1), aggrecan (Acan), and VEGF (Vegfa) (Table 1). The expression levels were
analyzed and compared to the housekeeping gene glyceraldehydes 3-phosphate
dehydrogenase (Gapdh). The specificity of the primers was confirmed separately
before the real-time PCR reaction. The expression of the tested genes was
calculated via the 2-∆∆Ct method [22] using HA/TCP as the reference group.
2.3.4 Histological analysis
All the samples were fixed in 10% phosphate buffered formalin for 1 hour before
embedding. The in vitro PLGA/PCL and collagen I scaffolds were embedded in
paraffin (n=3). The microtome sections with 6 μm of thickness were stained with
0.4% thionin (Sigma) after deparaffinization in xylene and rehydration through
graded ethanol. Simultaneously, HA/TCP and PU scaffolds were embedded in
methylmetacrylate (MMA, L.T.I., Bilthoven, the Netherlands) because of their rigid
character. The samples were sectioned to 15μm using a microtome equipped with
diamond blade (Leica SP1600, Leica microsystem, Wetzlar, Germany) and then
stained with 0.4% thionin solution.
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Performance of Different 3D Scaffolds for Endochondral Bone Formation
2.4 In vivo implantation
After 4 weeks of in vitro culture, the cell-scaffold constructs were implanted
subcutaneously in ten 8-week-old male nude rats (Crl:NIH-Foxn1mu, Charles River).
The protocol was approved by the Animal Ethical Committee of the Radboud
University Nijmegen Medical Centre (Approval no: RU-DEC 2011-142) and the
national guidelines for the care and use of laboratory animals were applied. All
rats received analgesic before and after the surgery. The surgical operations were
performed under general inhalation anesthesia with a combination of isoflurane,
nitrous oxide, and oxygen. The back of the rats was shaved and disinfected with
povidone-iodine. Subsequently, four small longitudinal incisions were made.
Lateral to each incision, a subcutaneous pocket was created. Each rat received
a cell-scaffold construct from each of the four materials and the constructs were
randomized. After placing the samples, the skin was closed using staples. After 8
weeks, the rats were euthanized by CO2-suffocation for sample collection.
2.5 Histological analysis of the in vivo samples
All in vivo samples were fixed in 10% phosphate buffered formalin for 30 hours and
dehydrated through graded ethanol before embedding either in MMA or paraffin.
MMA embedded samples (n = 5) were sectioned to 15μm using a microtome with
diamond blade (Leica SP1600, Leica microsystem, Wetzlar, Germany) and then
stained with methylene blue and basic fuchsin (both reagents from Merck).
Prior to paraffin embedding, the fixed samples (n = 5) were decalcified by 10%
EDTA and the decalcification process was monitored via X-ray. Five μm-thick
sections were made from each sample and used for hematoxylin/eosin (HE),
collagen II and tartrate-resistant acid phosphatase (TRAP) staining. For collagen
II staining, sections were treated with 0.1% pronase and 1% hyaluronidase for
antigen retrieval (both from Sigma). Afterwards the sections were incubated for 1
h at room temperature with mouse monoclonal antibody against collagen type II
(II-II6B3 antibody, 1:100; Developmental Studies Hybridoma Bank, Iowa City, USA)
linked with a biotin-labeled secondary antibody (1:500, Jackson Immuno Research,
Newmarket, England). Then alkaline phosphatase-conjugated streptavidin
(1:50, Biogenex HK-321-UK) was used as the third antibody and the activity was
visualized by using a new fuchsin substrate (Sigma). An isotype immuno-globulin
G1 monoclonal antibody was used as negative control. Counterstaining was
performed with Gill’s hematoxylin (Sigma). For TRAP staining, sections were first
incubated in TRIS/MgCl2-buffer for 1 hour. Afterwards they were stained with acid
phosphatase substrate solution at 37 °C for 30 min, which contained Naphthol
AS-BI phosphate, sodium nitrite, N, N- dimethylformamide, potassium sodium
tartrate and pararosaniline (all reagents from sigma).
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2.6 Polarized light microscopy
The in vivo MMA cross-sections were observed under polarized light using a
light microscope (DM RBE Leica, Bensheim, Germany) in which the polarizer and
analyzer were fixed perpendicularly to each other (cross-polarized light).
2.7 Statistic analysis
The data were analyzed using one-way ANOVA followed by Tukey’s Multiple
Comparison Test (GraphPad Prism 5, GraphPad Software Inc., San Diego, USA).
Results are reported as mean values and standard deviation. Differences were
considered statistically significant at p < 0.05.
3. Results
3.1 Cell seeding efficiency
After 24 hours of cell seeding, DNA content of the seeded cells on each group was
measured to assess the seeding efficiency (Fig. 2a). Highest cell number was found
on collagen I group, as all of the cells were encapsulated in the gel during cell
seeding. Comparable amount of cells was detected on PU and PLGA/PCL scaffolds,
which was around 65% of the total cell number seeded onto the scaffolds. In
comparison, lowest cell number was shown on HA/TCP scaffolds, with the seeding
efficiency of around 25%.
3.2 DNA and GAG content pre-implantation
After 1 week of proliferation and 4 weeks of chondrogenic differentiation in vitro,
comparable amounts of DNA could be found in HA/TCP, PLGA/PCL and collagen
Figure 2. DNA (a and b) and GAG (c) content on different scaffolds. DNA content after
24 hours of seeding (a) was measured to evaluate the seeding efficiency. GAG content
(c) was measured after 4 weeks of chondrogenic priming to assess the cartilage
formation and normalized by the amount of DNA (b). ** indicates statistically
significant difference at p<0.01; error bars represent standard deviation (n=3).
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Performance of Different 3D Scaffolds for Endochondral Bone Formation
I scaffolds, which was significant higher than that on the PU scaffolds as shown
in Fig. 2b. Likewise, the normalized GAG content on the aforementioned three
scaffolds was also comparable, but the PLGA/PCL constructs showed significant
higher GAG/DNA ratios than the PU samples (Fig. 2c).
5
Figure 3. Gene expression of aggrecan (a), collagen II (b), collagen X (c) and VEGF (d)
after in vitro chondrogenic priming. The gene expression level of the cells seeded on
HA/TCP scaffold was used for normalization. * indicates statistically significant
difference at p<0.05; ** indicates statistically significant difference at p<0.01; error
bars represent standard deviation (n=4). The red dash line indicates the gene
expression level of the reference group (HA/TCP).
3.3 Expression of chondrogenic genes pre-implantation
After 4 weeks of chondrogenic differentiation in vitro, a 2-fold higher level of
aggrecan expression was observed for HA/TCP group compared to collagen I
group, while there was no difference detected among HA/TCP, PU and PLGA/PCL
groups (Fig. 3a). Similar result was also found for VEGF expression that HA/TCP
group showed significantly higher level of VEGF compared to collagen I group (Fig.
3d). For collagen II gene expression, all groups showed a similar level (Fig. 3b). In
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contrast, a significant higher expression of collagen X was detected for PLGA/PCL
groups (Fig. 3c).
Figure 4. In vitro chondrogenic differatiation. The cartilage matrix deposition (stain
purple) in the scaffolds was visualized by thionin staining after 4 weeks of in vitro
culture. On HA/TCP (a and e) and PU (b and f), the cartilage matrix was formed in the
pores of the scaffolds. The cells exhibited typical chondrocyte like morphology and
hypertrophy could be observed, whereby the lacunae of the cells were enlarged and
the matrix was reduced (e and f). The shrinkage of the PLGA/PCL (c and g) and
collagen I (d and h) scaffolds during the histological processing compromised the
morphology of the cells. However, obvious cartilage matrix deposition was displayed
in both scaffolds and a homogeneous distribution of cells could be observed. (S:
scaffolds)
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Performance of Different 3D Scaffolds for Endochondral Bone Formation
3.4 Histological evaluation pre-implantation
To visualize the cartilage matrix deposition in the scaffolds, histological sections
were made and stained with thionin after 4 weeks of chondrogenic culture. For
both HA/TCP and PU, as shown in Fig. 4a and b, the cartilage matrix was formed
in the pores of the scaffolds and stained purple. The cells had a rounded or
polygonal shape and formed isogenous groups, exhibiting typical chondrocyte like
morphology (Fig. 4e and f). Furthermore, the majority of the cells had reached
the hypertrophic stage whereby the lacunae were enlarged and the matrix was
reduced.
The shrinkage of the fibrous scaffold during the histological processing compromised
the morphology of the cells on PLGA/PCL. Nevertheless, a substantial GAG content
was displayed in the scaffolds stained strongly in purple, representing the cartilage
matrix (Fig. 4c). The PLGA/PCL fibers were embedded within the matrix and
a homogeneous distribution of cells could be observed throughout the fibrous
structure (Fig. 4g). Similarly, the collagen I scaffolds also showed cartilage matrix
deposition although here the shrinkage was less, while the total amount appeared
to be lower than PLGA/PCL scaffolds. (Fig. 4d and h).
3.5 Histological analysis of in vivo bone formation
All animals recovered uneventfully from the surgical procedure and remained in
good health. No signs of wound complications were observed post-operatively.
After 8 weeks of implantation, all samples were retrieved and an obvious increase
in hardness of PLGA/PCL, collagen I and PU scaffolds could be sensed by hand.
Macroscopic signs of inflammation or adverse tissue responses were absent for all
retrieved samples.
To evaluate the bone forming quality and quantity, different histological staining
methods were used for either MMA (Fig. 5) or paraffin (Fig.6) embedded samples.
Bone formation was found in all HA/TCP, PU and collagen I constructs, and nine
(out of ten) PLGA/PCL scaffolds. For the bone-formed samples within each group,
the bone forming quality was similar as observed from the histological stainings.
Due to the deformation of PU and PLGA/PCL scaffolds during the tissue processing
procedures, the original material area could not be defined to quantify the ratio of
bone formation within the scaffolds.
3.5.1 HA/TCP scaffolds
As shown in Fig. 5a, substantial bone formation was present at both the periphery
and inside of the scaffolds. Within the macro-pores, the structure of a marrow
cavity filled with hematopoietic cells and adipocytes could be observed, surrounded
by concentric deposition of bone matrix. Moreover, plenty of blood vessels
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penetrated randomly in these bone forming areas (Fig. 6a and b). Osteocytes
were embedded in the bone matrix, showing relatively round shapes. Collagen II
staining confirmed that only a small amount of cartilage (stained pink) remained,
which was occasionally found in the small pores of a few samples where cells kept
the chondrocyte like shape without breaking down the surrounding matrix (Fig.
6c).
Figure 5. Overview of in vivo bone formation. MMA sections of HA/TCP (a), PU (b),
PLGA/PCL (c) and Collagen I (d) samples after 8 weeks of implantation were stained
with methylene blue and basic fuchsin. Bone formation (stained red) displays in
different spatial arrangements on each type of scaffold.
3.5.2 PU scaffolds
PU scaffold underwent deformation during the histological processing. Bone
formation was observed in the scaffolds, especially in the pores located in the
peripheral area (Fig. 5b). Surrounded by the bone tissue, a marrow-like structure
or a loosely organized fibrous tissue-like structure was detected in these pores
together with blood vessel in-growth (Fig. 6d and e). The amount of bone formation
in the central pores was relatively low compared to HA/TCP scaffolds. Chondrocytes
or cartilage like structure could not be seen at this time point. Collagen II staining
revealed a slight amount of tissue stained light pink, indicating the replacement of
in vitro cartilage with bone tissue was nearly completed (Fig. 6f).
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Performance of Different 3D Scaffolds for Endochondral Bone Formation
3.5.3 PLGA/PCL scaffolds
Extensive cartilage remodeling could be observed inside of PLGA/PCL scaffolds.
Islets of bone tissue were scattered amongst the remnants of scaffold material,
5
Figure 6. Histological analysis of in vivo bone formation on scaffolds. After 8 weeks of
implantation, HE (left and middle column) staining showed that the bone formation
on HA/TCP (a), PU (d) and Collagen I (j) scaffolds was associated with bone marrow
development while on PLGA/PCL scaffold (g) the bone tissue displayed a woven bone
like structure without concentrically organized structure and marrow formation. In
the magnified images of the black squared box (middle column), osteocytes were
embedded in the bone matrix and blood vessels penetrated in the bone forming areas
on all scaffolds. Collagen II staining (right column) confirmed significant remaining of
cartilage matrix (red color) on PLGA/PCL (i) while the remodeling of in vitro cartilage
into bone tissue was nearly completed on HA/TCP (c), PU (f) and Collagen I (l)
scaffolds. The black arrow indicates a blood vessel and the arrow head in the white
magnified box indicates embedded electrospun fibers. (B: bone; BM: bone marrow; S:
scaffold; C: cartilage)
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indicating the onset of the newly formed bone. Concentrically organized bone
tissue with marrow cavity structures could not be observed (Fig. 5c). HE staining
Figure 7. Histological analysis of in vivo osteoclast activity. After 8 weeks of in vivo
implantation, the activity of osteoclasts (stain red) was visualized by TRAP staining.
For HA/TCP, PU and collagen I scaffolds, osteoclast activity was positive (Fig. 7a, b
and d) which scattered along the edge of bone matrix (Fig. 7e, f and h). In comparison,
an extensive level of osteoclast activity was observed throughout the PLGA/PCL
scaffolds (Fig. 7c), which was not only located along the edge of new bone formation,
but also present in the middle area of the bone/cartilage mixed matrix (Fig. 7g).
96
Performance of Different 3D Scaffolds for Endochondral Bone Formation
revealed that the bone tissue displayed a woven bone like structure with vascular
invasion. A large number of immature osteocytes were embedded in the matrix
and randomly arranged (Fig. 6g and h). Interestingly, the remained PLGA/PCL
fibers of the scaffolds were also embedded in the matrix as indicated in Fig. 6h.
Furthermore, the collagen II staining (Fig. 6i) confirmed the existence of the
remaining cartilage matrix (red color). Compared to the PLGA/PCL constructs
before implantation (Fig. 3c), the in vivo samples contained considerably reduced
amount of cartilage matrix. However, the newly formed bone still resembled a
mixture of cartilage and calcified cartilage tissue or bone.
3.5.4 Collagen scaffolds
Bone formation with marrow development was observed in collagen I gel. An
intact bone layer was displayed along the fibrous capsule. Trabecular bone-like
structure, together with a great number of adipocytes and hematopoietic cells,
was distributed in the middle part (Fig. 5d). Blood vessels penetrated in the stroma
and were also present inside the bone tissue (Fig. 6j and k). There was no remaining
cartilage matrix visualized by a collagen II staining which indicated the complete
replacement of in vitro cartilage by new bone tissue (Fig. 6l).
3.6 Histological analysis of osteoclast activity
The activity of osteoclasts was visualized by TRAP staining. After 8 weeks of in
vivo implantation, osteoclast activity was positive on HA/TCP, PU and collagen I
scaffolds (Fig. 7), which scattered along the edges of bone matrix. In comparison,
an extensive level of osteoclast activity was observed throughout the PLGA/PCL
scaffolds, which was not only located along the edges of new bone formation, but
also present in the middle area of the bone/cartilage mixed matrix.
3.7 Polarized light observation
The collagen bundle arrangement of the newly formed bone tissue on each scaffold
was examined under the polarized light on the basis of its birefringent character.
As shown in Fig. 8, the collagen bundles within the bone forming areas which were
oriented transversely to the direction of the light propagation were visualized in
bright pink under the polarized light. In HA/TCP (Fig. 8a and e), PU (Fig. 8b and
f) and collagen I scaffolds (Fig. 8d and h), a parallel pattern of the collagen fibril
arrangement could be observed partially in the bone forming regions, indicating
the initial organization of the lamellar bone. However, in the PLGA/PCL scaffolds
(Fig. 8c and g), the bone matrix displayed a random spatial organization, as shown
by an isotropic and un-polarized structure.
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Figure 8. Observation of in vivo bone formation under polarized light. Newly formed
bone tissue in each scaffold (embedded in MMA) was stained with methylene blue /
basic fuchsin (a-d) and examined under the polarized light (e-h). The bone tissue was
stained red (a-d) and the collagen bundles which were oriented transversely to the
direction of the light propagation were distinguished in bright pink under the
polarized light (e-h). In HA/TCP (a and e), PU (b and f) and collagen I (d and h)
scaffolds, a concentric and parallel pattern of the collagen fibril arrangement could
be observed partially in the bone forming regions. For PLGA/PCL scaffolds (c and g),
the bone matrix displayed an isotropic organization. (Scale bar: 50µm)
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Performance of Different 3D Scaffolds for Endochondral Bone Formation
4. Discussion
An appropriate scaffold material is an essential component in bone tissue
engineering as it regulates the cellular behaviors, such as cell proliferation and
differentiation, on both chondrogenic and osteogenic pathways [23, 24]. In this
study, we aimed to evaluate and compare the efficacy of four scaffolds for bone
tissue regeneration via the endochondral ossification, which were porous HA/TCP
composites, porous PU foam, PLGA/PCL electrospun fibers and collagen I gel.
Our results showed different cell seeding efficiencies on the tested scaffolds due to
their varied structural features, among which the HA/TCP scaffolds had the lowest
initial cell number. Nevertheless, 1-week culture in proliferation medium ensured
a comparable amount of cells present on these scaffolds prior to the implantation,
as the low-serum chondrogenic culture condition would principally not stimulate
the cell proliferation. The results also indicated that, all of the selected scaffolds
supported cell growth and proliferation. Similarly, comparable cartilage formation
in the scaffolds was also shown before the in vivo implantation. However, cells
cultured on the PLGA/PCL scaffolds expressed a significantly higher level of collagen
X compared to those on the other types of scaffolds. In vivo results revealed that,
8 weeks of implantation was sufficient to ossify nearly the complete constructs of
HA/TCP, PU and collagen constructs and form a hematopoietic bone marrow, while
the bone formation on PLGA/PCL scaffolds displayed a immature structure with a
random spatial organization of collagen bundle and obvious cartilage remaining.
Furthermore, an extensive level of osteoclast activity present throughout the
PLGA/PCL scaffold also demonstrated a highly active cartilage and bone remodeling
in PLGA/PCL compared to the other three scaffolds. These results indicated that,
although the cells on PLGA/PCL might have an earlier hypertrophic stage in vitro,
the cartilage remodeling and vessel invading in PLGA/PCL was less efficient in vivo
compared to the other scaffolds.
Hypertrophic cartilage remodeling is an critical process to initialize the successive
endochondral bone formation and such progression is closely related to the
reciprocal interaction between the cells and natural ECM [25]. Cellular products, e.g.
proteinases and VEGF, modify the type X collagen-rich ECM and promote vessel
formation; ECM also releases sequestered growth factors and cytokines to regulate
the endochondral commitment of the cells [26]. Hypertrophic cartilage remodeling
usually couples the expression of VEGF [27]. Surprisingly, a significantly higher
level of collagen X expression on PLGA/PCL scaffolds did not result in an elevated
VEGF expression of the cells. This may consequently lead to the formation of less
mature bone in PLGA/PCL scaffolds compared to the other scaffolds. Moreover, a
comparable endochondral bone formation was observed on HA/TCP and collagen
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I gel, although significant difference of VEGF expression was shown for these two
scaffolds prior to the implantation. Therefore, apart from the natural ECM, the
scaffolds, acting as a synthetic ECM, also played an essential role in directing the
cells to form endochondral bone in vivo.
Noticeably, our data demonstrated that the structure and maturity of the
endochondral bone is material/scaffold dependent. Among various scaffold
properties, geometry and composition played important roles. In vivo, higher
percentage of porosity allows for better cell recruitment and vascularization,
which leads to improved osteogenesis [28]. On top of that, scaffold pore size can
dictate the structure and the mode of the newly formed bone (intramembranous
or endochondral) [24]. In an in vivo study of BMP-2 induced osteogenesis and
chondrogenesis via honey-comb-shaped HA scaffolds, small diameter ‘tunnels’ (90120µm) favored chondrogenesis followed by osteogenesis, while large diameter
‘tunnels’ (350µm) favored intramembranous bone formation [24, 29]. Regarding the
bone structure, increased trabeculae formation appeared to be supported when
the scaffolds have smaller pores [30]. Similarly, in our study, more lamellar-like bone
structure with bone marrow formation was generated on HA/TCP and PU scaffolds
compared to PLGA/PCL. This might be related to the larger pore size of HA/TCP and
PU scaffolds, which can facilitate more oxygen and nutrient diffusion and allow
easier vessel in-growth during cartilage/bone remodeling. In comparison, the
small pore size of PLGA/PCL scaffolds might have created a hypoxic environment
which favored the cartilage maintenance [28]. Furthermore, the confined porous
structure of HA/TCP and PU can concentrate collagen fibers within the pores
and triggered their self-assembly, thereby forming the lamellar like bone [31].
Conversely, in fibrous PLGA/PCL scaffolds, new collagen fibers formed by the cells
were distributed in a random manner, resulting mostly in woven bone [31]. Thus,
the specific design of the scaffold can drive the arrangement of collagen fibers and
consequently target bone formation.
Bone in-growth to the central part of the HA/TCP scaffold had a relatively higher
volume compared to the PU scaffolds, even though HA/TCP scaffolds were
a little thicker. The main reason for the difference of bone in-growth might be
that, unlike PU, HA and TCP are osteoconductive [32, 33]. This osteoconductive
property can promote the attachment of bone-forming cells from the host as
well as cell migration and vessel formation [14]. Both PU and HA/TCP are slowly
degrading materials which were unlikely to degrade after 8 weeks of implantation
[19, 34]
. However, compared to PU scaffolds, HA/TCP could actively release calcium
and phosphate ions which can influence the local ionic concentration to trigger
and provide the source for the deposition of calcium in the cartilage matrix for
subsequent remodeling [34]. With respect to PLGA/PCL scaffolds and collagen I gel,
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Performance of Different 3D Scaffolds for Endochondral Bone Formation
the naturally derived collagen I component resembles the organic composition of
natural bone and is considered more osteoconductive than the synthetic polymers,
i.e. PLGA and PCL [35], which have also contributed to the differences of the bone
formation between these two scaffolds.
For in vivo bone generation, the origin of bone cells and the ossification type are
closely related to the nature and commitment of the seeded cells. Previous results
have suggested a tendency of MSCs to undergo endochondral ossification when
implanted in vivo. Hartman et al. seeded rat MSCs onto porous HA/TCP scaffold to
investigate ectopic bone formation via the intramembranous approach, while the
onset of bone tissue was always surrounded by hypertrophic cartilage-like cells [36].
Tortelli et al. combined murine MSCs or mature osteoblasts with porous ceramic
scaffolds and subsequently implanted them ectopically in mice [37]. New vascularized
bone was formed through the activation of the endochondral ossification process
in the MSC-scaffold complex. Conversely, osteoblasts directly formed bone
via an intramembranous ossification, without any signs of vascularization [37].
Aforementioned results indicated that, activating MSCs toward an endochondral
ossification route is presumably in line with their natural developmental pattern
and provides significant advantages in generating vascularized bone tissue. Our
results demonstrated the tendency of in vitro chondrogenically induced MSC to
ossify in a non-chondrogenic environment in vivo, which is also consistent with
some previous studies [15, 38, 39]. When combined with various scaffolds, scaled-up
bone formation with the structure and functionality comparable to that of native
bones could be achieved in vivo without the use of any osteoinductive growth
factors, e.g. BMP-2, at the implant site. Among the tested scaffold materials, HA/
TCP block and collagen I gel supported most mature bone formation within the
entire constructs. Regarding the clinical applications, Collagen I gel is appealing
for irregularly shaped defects and its fast degradation allows rapid turnover of
the new bone tissue in the implantation site. HA/TCP scaffold is more suitable
for the implantation in the load-bearing areas, which can provide initial structural
stability. Nevertheless, considering the ceramic scaffolds are brittle, the mechanical
property of HA/TCP needs to be further modified, for instance by incorporating
polymeric components.
Some limitations of our study have to be addressed: i) the evaluation was only
performed at one time point of 8 weeks after implantation, ii) the scaffolds were
in a relatively small size which might favor cell invasion and vascularization, iii)
the implantation was only performed in a subcutaneous environment where the
mechanical stimuli are not involved, and iv) due to the differences in composition
and structure of the selected scaffolds, the specific parameters of a material
scaffold that dominate the endochondral bone forming process could not be
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concluded from the current study. More completed understanding will be obtained
if all structural parameters of the scaffolds are kept constant and meanwhile only
the composition is varied, or vice versa. Further investigations are warranted to
dynamically monitor the endochondral bone forming process in an orthotopic
implantation site. Furthermore, systematical investigations are also needed to
address how this process is driven by the parameters of the supporting materials,
so that more appropriate scaffolds can be selected and tailored to target largescale bone regeneration.
5. Conclusion
In this study, chondrogenic priming rat MSCs in the presence of different 3D
scaffold materials followed by subcutaneous implantation allowed us to establish
a reliable scaffold-based heterotopic model for endochondral bone formation.
The bone quality and maturity was scaffold/material-dependent. Eight weeks of
implantation was not sufficient to ossify the entire PLGA/PCL constructs, while a
comprehensive remodeling of the cartilage has occurred. HA/TCP, PU and collagen
I scaffolds supported full bone formation with rich vascularity and marrow stroma
development. This study revealed that bone could be generated heterotopically
via the endochondral pathway in absence of osteoinductive growth factors and
scaffold properties could dictate the morphology and maturity of endochondral
bone formation. Continued research in this direction would open the attractive
possibility to target large-scale bone regeneration.
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Chapter 6
Effects of in vitro chondrogenic
priming time of bone marrow derived
mesenchymal stem cells on in vivo
endochondral bone formation
Wanxun Yang, Sanne K. Both, Gerjo J. V. M.van Osch,
Yining Wang, John A. Jansen and Fang Yang
Effects of In Vitro Priming Time on Endochondral Bone Formation
Summary of this chapter
Recapitulation of endochondral ossification leads to a new concept of bone tissue
engineering via a cartilage intermediate as an osteoinductive template. To enable
an effective translation of this endochondral approach to clinic, it is imperative
to possibly shorten the in vitro procedure for the creation of cartilage template
and meanwhile ensure a sufficient bone formation. Therefore, in the chapter,
we aimed to investigate the influence of in vitro chondrogenic priming time on
the in vivo endochondral bone formation both qualitatively and quantitatively.
To this end, rat bone marrow derived mesenchymal stem cells were seeded
on two scaffolds with distinguished features, i.e. a fibrous poly(lactic-coglycolic acid)/poly(ε-caprolactone) electrospun scaffold (PLGA/PCL) or a porous
hydroxyapatite/tricalcium phosphate composite (HA/TCP). Next the constructs
were chondrogenically differentiated for 2, 3 and 4 weeks in vitro followed by
subcutaneous implantation for up to 8 weeks. A longer chondrogenic priming
time resulted in a significantly increased amount and homogeneous deposition
of the cartilage matrix on both PLGA/PCL and HA/TCP scaffolds in vitro. In vivo, all
implanted constructs gave rise to endochondral bone formation, whereas the bone
volume was not affected by a longer priming time. An unpolarized woven bone
like structure, with significant amounts of cartilage remaining, was generated in
fibrous PLGA/PCL scaffolds, while porous HA/TCP scaffolds supported progressive
lamellar-like bone formation with mature bone marrow development. These data
suggest that, by utilizing a chondrogenically differentiated MSC-scaffold construct
as cartilage template, the essential in vitro priming time could be shortened to
2 weeks to generate a substantial amount of vascularized endochondral bone in
vivo. The structure of the bone was controlled by the chemical and structural cues
provided by the scaffold design.
1. Introduction
The bone forming process during development and regeneration is governed by
intramembranous or endochondral mechanism, through which bone is generated
by direct osteoblastic differentiation or remodelling from a cartilage intermediate,
respectively [1-3]. Recapitulation of endochondral ossification leads to a new
route to engineer bone tissue, which generally involves two steps: 1) creating a
cartilage intermediate in vitro and 2) implanting the cartilage intermediate as an
osteoinductive template to transform into bone [4, 5]. Recent studies have shown the
superiority of generating vascularized bone tissue via the endochondral approach,
which has the potential to overcome the issue of insufficient vascularization faced
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by the conventional bone tissue engineering based on the intramembranous
ossification [6-9].
In general, to create the cartilage intermediate, mesenchymal stem cells (MSCs)
loaded in a three dimensional scaffold or in a cell pellet form are cultured in
chondrogenic conditions for a certain time period before in vivo implantation.
The obtained cartilage serves as a morphological template enabling blood vessel
invasion and endogenous MSCs recruitment for successive bone formation [10]. It
is clear that such a cartilage construct contains essential “biological instructions”
to initiate and regulate the endochondral bone forming process, and this largely
relies on the stage of the chondrogenic differentiation, as chondrocytes possess
specific metabolic features and secrete diverse biochemical cues when they are in
different developmental stages [9, 11]. A fine coordination between the progression
of the cartilage construct and endochondral transformation needs to be launched
by choosing an optimal time period for in vitro chondrogenic priming.
In literature, the protocol applied for chondrogenically differentiating MSCs was
typically aimed to reach the stage with a certain degree of hypertrophy before
implantation, which usually lasts for 4 to 5 weeks [4, 6, 9]. The hypothesis behind
this approach was that hypertrophy could ensure the progression of endochondral
differentiation and secretion of substantial amount of vascular endothelial growth
factor (VEGF), which might enable timely establishment of vascularization into the
implants [12]. However, this assumption was made mainly based on the use of cell
pellets, and it has not yet been confirmed [13, 14]. For the final application, it is more
relevant to elucidate the optimal in vitro priming time for the creation of cartilage
template when a scaffold material is involved.
From a clinical point of view, a long in vitro culture time associates with
considerably high treatment cost and regulatory burden, which hampers the
effective translation of this tissue engineering strategy to the clinics. Therefore, it
is imperative to seek the possibility of shortening this procedure and meanwhile
ensure a sufficient endochondral bone formation. In order to gain more insight on
the endochondral approach for future clinical applications, we aimed to investigate
the influence of chondrogenic priming time on the in vivo endochondral bone
formation both qualitatively and quantitatively. Furthermore, to examine if such
influence is universal or also depends on the properties of the applied scaffolds,
two representative scaffolds with distinguished characteristics were selected
as cell carrier, i.e. a polymer-based fibrous poly(lactic-co-glycolic acid)/poly(εcaprolactone) electrospun scaffold (PLGA/PCL) and a ceramic-based porous
hydroxyapatite/tricalcium phosphate composite (HA/TCP). Rat bone marrow
derived MSCs were first seeded on these two scaffolds and cultured in chondrogenic
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Effects of In Vitro Priming Time on Endochondral Bone Formation
condition for 2, 3 and 4 weeks. Afterwards, these cell-scaffold constructs were
implanted subcutaneously for up to 8 weeks in nude rats to assess bone formation
and cartilage remains by histological stainings and histomorphometric analysis.
2. Materials and Methods
2.1 Cell isolation, seeding and culture
MSCs were isolated from 6-week-old male Fischer rats (Charles River) after the
approval from Radboud University Animal Ethics Committee (Approval no: RUDEC 2012-195). Briefly, two femora of each rat were removed and the epiphyses
were cut off. Rat MSCs were flushed out from the remaining diaphyses using the
proliferation medium through a needle. The proliferation medium consisted of alpha
Minimal Essential Medium (aMEM; Gibco, Grand Island, USA) supplemented with
10% fetal bovine serum (FBS; Sigma, St. Louis, USA), 100 U/ml penicillin, 100μg/ml
streptomycin (Gibco), 50 μg/ml l-ascorbic acid (Sigma) and 10nM dexamethasone
(Sigma). The flush-out was cultured for 1 day in a humidified incubator (37 °C, 5%
CO2), after which the medium was refreshed to remove non-adherent cells. The
cells were cultured for an additional 4 days then they were detached using trypsin/
EDTA (0.25% w/v trypsin, 0.02% w/v EDTA; Sigma) and counted.
The cells were seeded in two types of scaffold: 1) Fibrous poly(lactic-co-glycolic
acid) (PLGA, Purasorb® PDLG 8531, Purac Biomaterials BV, Gorinchem, the
Netherlands) combined with poly(ε-caprolactone) (PCL, LACTEL® Absorbable
Polymers, DURECT Corporation, Pelham, USA) electrospun scaffold made by a wetelectrospinning technique (Esprayer ES-2000S, Fuence Co., Ltd, Tokyo, Japan) as
previously described [7]. This PLGA/PCL scaffold was in a disk shape with 6 mm of
diameter and 2 mm of thickness. 2) Porous sintered hydroxyapatite (HA)/tricalcium
phosphate (TCP) composite (CAM biceramics, Leiden, the Nethelands) in a cubic
shape of 3×3×3 mm. This HA/TCP scaffold consisted of HA and TCP with a 75/25
weight ratio and the volumetric porosity was 75% (provided by the manufacturer).
Before cell seeding, the scaffolds were sterilized by either autoclave (HA/TCP) or in
70% ethanol for 2 hours (PLGA/PCL) and subsequently soaked in the proliferation
medium overnight after being washed with PBS. During cell loading, the scaffolds
were incubated in the cell suspension with a concentration of 1 × 106 cells/ml (4
scaffolds per 1 ml of cell suspension) and gently rotated for 3 hours. Afterwards the
unattached cells from the suspension were collected, centrifuged and reseeded
onto the scaffolds to ensure a high cell loading efficiency.
All cell-scaffold constructs were cultured for 1 week in proliferation medium.
Thereafter, the medium was changed to chondrogenic medium, consisting of
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high-glucose DMEM (Gibco), 1% FBS, 100 U/ml penicillin, 100 μg/ml streptomycin,
50 μg/ml l-ascorbic acid, 100 mM of sodium pyruvate (Gibco), 40μg/ml l-proline
(Sigma), 1:100 insulin-transferrin-selenium (ITS; BD Biosciences, Bedford, USA),
100nM dexamethasone, 10 ng/ml of transforming growth factor beta-2 (TGF-β2;
R&D systems, Minneapolis, USA) and 100 ng/ml bone morphogenetic protein
2 (BMP-2; BD biosciences). The constructs were cultured for 2, 3 or 4 weeks in
chondrogenic medium which was refreshed twice a week. At each time point, the
cell-scaffold constructs were collected and divided into two groups. One group of
samples were used for in vitro characterization and the other for subcutaneous
implantation.
2.2 In vitro tests
2.2.1 GAG content
The amount of GAG in each sample was detected by a sulfated glycosaminoglycan
(sGAG) assay (n=3). After being washed twice in PBS, all samples were digested
in proteinase K solution (1 mg/ml proteinase K in 50 mM Tris with 1 mM EDTA, 1
mM iodoacetamide, and 10 μg/ml pepstatin A) (all reagents from Sigma) for 20
hours at 56°C. The GAG content was determined by using dimethylmethylene blue
(Polysciences Inc., Warrington, USA), with chondroitin sulfate (Sigma) as standard
control. Measurements of absorption were performed at 530 nm and 570 nm
using an ELISA reader (Bio-Tek Instruments Inc., Winooski, USA). The results were
calculated according to the previously described method [15].
2.2.2 DNA content
DNA content (n=3) was quantified by QuantiFluorTM DNA system (Promega
Corporation, Madison, USA) using the same cell extract solution as used for
the GAG assay. Followed the manufacturer’s instruction, a DNA standard curve
was used to quantify the amount of DNA in each sample and the results were
measured using a fluorescence microplate reader (Bio-Tek Instruments Inc.) with
an excitation wavelength at 485nm and an emission wavelength at 530nm.
2.2.3 RNA extraction and qPCR analysis
After 2, 3 and 4 weeks of chondrogenic culture, RNA of the seeded cells was
extracted using TRIzol® reagent according to the manufacturer’s instructions (Life
Technologies, Carlsbad, USA). The obtained RNA was dissolved in RNase free water
and the concentration was measured with spectrophotometer (NanoDrop 2000,
Thermal scientific, Wilmington, USA). Subsequently, first strand cDNA was reverse
transcribed from RNA using iScript™ cDNA Synthesis Kit (Bio-Rad Laboratories, Inc.,
USA). Afterwards cDNA was further amplified and the expression of specific genes
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Effects of In Vitro Priming Time on Endochondral Bone Formation
was quantified using IQ SYBR Green Supermix PCR kit (BioRad, Hemel Hempstead,
United Kingdom) in a real-time PCR (BioRad, CFX96™ real-time system) (n=4).
Chondrogenic-related gene markers were evaluated, including collagen type II
(Col2α1), aggrecan (Acan), collagen type X (Col10α1) and VEGF (Vegfa) (Table 1).
The specificity of the primers was confirmed separately before the real-time PCR
reaction. The expression levels were compared to the housekeeping gene β-Actin
and calculated via the 2-∆∆Ct method [16].
Table 1. Primer sequences used for qPCR
Forward (5’ → 3’)
Reverse (5’ → 3’)
Col2α1
GACGCCACGCTCAAGTCGCT
CGCTGGGTTGGGGTAGACGC
Col10α1
GGGCCCTATTGGACCACCAGGTA
CCGGCATGCCTGTTACCCCC
Acan
CATTCGCACGGGAGCAGCCA
TGGGGTCCGTGGGCTCACAA
Vegfa
ACTCATCAGCCAGGGAGTCT
GGGAGTGAAGGAGCAACCTC
β-Actin
TTCAACACCCCAGCCATGT
TGTGGTACGACCAGAGGCATAC
2.2.4 Histological analysis
The in vitro HA/TCP scaffolds were fixed in 10% phosphate buffered formalin for 1
hour and embedded in methylmetacrylate (MMA, L.T.I., Bilthoven, the Netherlands)
(n=3). Sections of to 15μm were made using a microtome equipped with diamond
blade (Leica SP1600, Leica microsystem, Wetzlar, Germany) and then stained with
0.04% thionin solution (Sigma). The in vitro PLGA/PCL scaffolds were fixed in 10%
phosphate buffered formalin for 1 hour and embedded in paraffin (n=3). The
microtome sections with 5 μm of thickness were also stained with 0.04% thionin
solution after deparaffinization in xylene and rehydration through graded ethanol.
2.3 In vivo implantation
After 2, 3 and 4 weeks of in vitro culture, the cell-scaffold constructs were implanted
subcutaneously in 8-week-old male nude rats (Crl:NIH-Foxn1mu, Charles River).
The protocol was approved by the Animal Ethical Committee of the Radboud
University Medical Centre (Approval no: RU-DEC 2012-195) and the national
guidelines for the care and use of laboratory animals were applied. For each time
condition, 3 rats were used for the implantation. All rats received analgesic before
and after the surgery. Surgery was performed under general inhalation anesthesia
with a combination of isoflurane, nitrous oxide, and oxygen. The back of the
rats was shaved and disinfected with povidone-iodine. Subsequently, four small
longitudinal incisions were made and a subcutaneous pocket was created lateral
to each incision. Two HA/TCP and two PLGA/PCL constructs were implanted to
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each rat (one construct per pocket). After placing the samples, the skin was closed
using staples. After 4, 6 and 8 weeks, rats were euthanized by CO2-suffocation for
sample collection.
2.4 Histological staining of the in vivo samples
All in vivo samples were fixed in 10% phosphate buffered formalin for 30 hours
and dehydrated through graded ethanol. For each rat, half of the samples (one
HA/TCP and one PLGA/PCL construct) were embedded in MMA, and the other half
samples were embedded in paraffin.
MMA embedded samples (n = 3 for each time condition) were sectioned using
a microtome equipped with diamond saw (Leica SP1600, Leica microsystem,
Wetzlar, Germany) and then stained with methylene blue and basic fuchsin (both
reagents from Merck).
Prior to paraffin embedding, the fixed samples (n = 3 for each time condition) were
decalcified by 10% EDTA and the decalcification process was monitored via X-ray.
Five μm-thick sections were made from each sample and used for hematoxylin/
eosin (HE), Elastic Van Gieson (EVG) and collagen II staining.
2.5 Histological and histomorphometrical analysis
Digital images of all histological sections were made using a light microscope
equipped with a digital camera (Axio Imager Z1, Carl Zeiss AG Light Microscopy,
Göttingen, Germany). Histological analysis consisted of a concise morphological
and histological evaluation of all sections. Histomorphometrical analysis was
performed with computer-based quantitative image analysis techniques based on
color recognition with manual correction (Leica QWin Pro®, Leica Microsystems
AG, Wetzlar, Germany). For both HA/TCP and PLGA/PCL scaffolds, quantification
of bone and cartilage area was performed on EVG staining images and collagen
II staining images, respectively. The region of interest (ROI) in each image
was defined from the total area of the cross section (including the scaffold
region) which was manually set using Leica QWin software. Bone and cartilage
tissue were discriminated by the difference in staining intensity and cellular
morphology. The parameters “bone area” and “cartilage area” were evaluated and
defined as a percentage of bone or cartilage tissue within the ROI, respectively.
Histomorphometrical data were collected based on three sections per specimen
and averaged to one value per specimen.
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Effects of In Vitro Priming Time on Endochondral Bone Formation
2.6 Statistic analysis
The data was analyzed using two-way ANOVA followed by Bonferroni’s Multiple
Comparison Test (GraphPad Prism 5, GraphPad Software Inc., San Diego, USA).
Mean values and standard deviations (SD) are reported in each figure as well
as relevant statistical relationships. Differences were considered statistically
significant at p < 0.05.
3. Results
3.1 In vitro cartilage formation pre-implantation
3.1.1 GAG production and DNA content
As shown in Fig. 1a, on both PLGA/PCL and HA/TCP scaffolds the total amount
of GAG deposition accumulated with chondrogenic culture time and reached the
highest amount in week 4. PLGA/PCL demonstrated significant higher amount
6
Figure 1. GAG (a) and DNA (b) content of PLGA/PCL and HA/TCP constructs after in
vitro chondrogenic priming. * p<0.05, ** p<0.01; error bars represent standard
deviation of the mean (n=3)
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of GAG compared to that on the HA/TCP scaffolds. Furthermore, a comparable
DNA amount was detected on both PLGA/PCL and HA/TCP, which remained stable
during the entire culture period (Fig. 1b).
Figure 2. Gene expression of collagen II (a), aggrecan (b), collagen X (c) and VEGF (d)
after in vitro chondrogenic priming. * p<0.05, ** p<0.01; error bars represent
standard deviation of the mean (n=4)
3.1.2 Expression of chondrogenic genes
For both PLGA/PCL and HA/TCP scaffolds, Col2α1 and Acan expression mantained
stable during the in vitro chondrogenic priming process and comparable expression
levels were detected on both scaffolds (Fig. 2a and b). On the PLGA/PCL constructs
there was a peak for the Col10α1 expression at week 3. In comparison, for the
HA/TCP constructs, the expression of Col10α1 remained constant till week 3 and
then decreased dramatically at week 4. Moreover, significantly higher expression
of Col10α1 was shown on the PLGA/PCL constructs compared to that on HA/TCP
after 3 or 4 weeks of chondrogenic priming (Fig. 2c). Regarding the expression
of Vegfa, HA/TCP constructs demonstrated a significantly higher level than that
of PLGA/PCL. Overall, for both scaffolds, the level of Vegfa expression remained
constant during the in vitro priming.
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Effects of In Vitro Priming Time on Endochondral Bone Formation
6
Figure 3. In vitro chondrogenic differatiation in PLGA/PCL and HA/TCP scaffolds after
2 and 4 weeks of culture. Cartilage matrix formation (stained purple) and cell
penetration inside of the PLGA/PCL scaffolds could be observed after 2 weeks of
chondrogenic priming (a and c). Nevertheless, the cartilage matrix mainly located on
one side of the scaffolds (a). When the culture time reached 4 weeks, the cartilage
matrix spread throughout the entire fibrous structure (b) and a great number of cells
with typical chondrocyte-like morphology could be observed (d). For HA/TCP scaffolds,
the cartilage formation could be observed on the pore surfaces of 2-week primed
constructs (e), while the cell penetration and matrix deposition were very limited in
the center of the scaffolds (e). Chondrocyte-like cells were embedded in the matrix
and mixed with undifferentiated cells (g). As the priming time increased, the cartilage
matrix deposition was augmented and distributed more homogeneously in the
scaffolds (f). After 4 weeks, the presented cells in the scaffolds mainly consisted of
chondrocyte-like cells (h).
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3.1.3 Histological analysis
To visualize the cartilage matrix deposition in the scaffolds, sections were made and
stained with thionin. During the histological processing, PLGA/PCL scaffolds shrank
and formed folds on the surfaces as shown in Fig. 4a. Yet cartilage matrix (GAG)
formation and cell penetration inside of the scaffolds could be clearly observed
after 2 weeks of chondrogenic priming (Fig. 3a and c). Nevertheless, the cartilage
matrix mainly located on one side of the scaffolds. When the culture time reached
4 weeks, the cartilage matrix spread throughout the entire fibrous structure (Fig.
3b) and a large number of cells with typical chondrocyte-like morphology could be
observed (Fig. 3d).
For HA/TCP scaffolds, the cartilage formation could be observed on the pore surfaces
after 2 weeks of chondrogenic priming, while the cell penetration and matrix
deposition were limited in the center of the scaffolds (Fig. 3e). Typical chondrocytelike cells were embedded in the matrix and mixed with undifferentiated cells (Fig.
3g). As the chondrogenic priming time increased, the cartilage matrix deposition
was augmented and distributed more homogeneously in the scaffolds (Fig. 3f).
After 4 weeks, the cells present in the scaffolds mainly consisted of chondrocytes
which were surrounded by abundant matrix and appeared in isogenous groups or
with large lacunae (Fig. 3h).
3.2 In vivo bone formation
3.2.1 General observations
From the total 27 animals available for surgery, one animal died during anesthesia.
Due to this, 3-week in vitro priming group had one animal less than the other
groups for evaluating the bone formation at the 8-week time point. The remaining
26 animals recovered uneventfully from the surgical procedure and remained in
good health. No signs of wound complications were observed post-operatively. At
the time point of 4, 6 or 8 weeks, all samples were retrieved and macroscopic signs
of inflammation or adverse tissue responses were absent for all of the retrieved
samples. 3.2.2 Descriptive histology
Bone formation was found in all PLGA/PCL and HA/TCP scaffolds. The bone forming
quality and quantity were similar in all samples within the same group.
3.2.2.1 PLGA/PCL scaffold
With different in vitro chondrogenic priming time, extensive cartilage remodeling
has already occurred within the PLGA/PCL scaffolds after 4 weeks of implantation
(Fig. 4a and b). The bone/cartilage complex distributed in the scaffolds in a
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Effects of In Vitro Priming Time on Endochondral Bone Formation
dispersive manner and displayed a trabecula-like structure. The implanted scaffolds
that received 4 weeks of chondrogenic priming in vitro showed higher amount
of bone/cartilage matrix presence in vivo with less fibrous tissue occupation (Fig.
4b) compared to the constructs which were chondrogenically differentiated for 2
weeks (Fig. 4a). From the HE staining pictures, a mixture of bone and cartilage tissue
could be observed in all PLGA/PCL scaffolds independent of in vitro priming and
in vivo implantation time (Fig. 5). Bone matrix was frequently found deposited on
the periphery of cartilage tissue or around the enlarged lacunae of chondrocytes.
The newly formed bone tissue resembled a woven bone structure, in the
surrounding of which the blood vessel formation could occasionally be observed.
Concentrically organized bone tissue and marrow development could not be seen
(Fig. 5). Furthermore, the non-degraded electrospun fibers of the scaffolds were
also integrated with the bone/cartilage matrix and surrounding tissue.
6
Figure 4. Overview of in vivo bone formation. MMA sections of PLGA/PCL (a and b)
and HA/TCP (c and d) constructs after 4 weeks of implantation were stained with
methylene blue and basic fuchsin. In PLGA/PCL constructs, the bone/cartilage
complex distributed in the scaffolds in a dispersive manner and displayed a trabeculalike structure. With 2 weeks of in vitro chondrogenic priming time, extensive cartilage
remodeling has already occurred (a), while 4-week primed PLGA/PCL constructs
showed higher amount of bone/cartilage matrix presence in vivo with less fibrous
tissue occupation (b). For HA/TCP scaffolds, bone formation occurred throughout the
whole implants with 2-week (c) or 4-week (d) in vitro priming, which was characterized
by close apposition of matrix on the surface of the pores together with islet or bridge
like bone tissue in the middle space of the pores.
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Chapter 6
Figure 5. Morphology of in vivo bone formation in PLGA/PCL constructs (HE staining).
A mixture of bone and cartilage tissue could be observed in all PLGA/PCL scaffolds
independent of in vitro priming and in vivo implantation time. Bone matrix (stained
red) was frequently found deposited on the periphery of cartilage tissue (stained
purple) or around the enlarged lacunae of chondrocytes. The newly formed bone
resembled a structure of the woven bone with vascular infiltration in the surrounding
tissue. Concentrically organized bone tissue and marrow development could not be
seen. (B: bone; C: cartilage.)
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Effects of In Vitro Priming Time on Endochondral Bone Formation
6
Figure 6. Morphology of in vivo bone formation in HA/TCP constructs (HE staining).
The length of the in vitro priming time did not show obvious influences on the
structure of the newly formed bone (stained red) at each in vivo time point. After 4
weeks of in vivo implantation, the new bone displayed in an irregular shape and the
osteoblasts could be recognized aligning at the periphery of the bone matrix (a and
e). The cartilage remnants (stained purple) were observed to fill small pores or mix
with the bone tissue (b and f). Marrow stroma development was initiated and filled
with abundant hematopoietic cells. After 8 weeks in vivo, the matrix exhibited a
concentric pattern and the collagen bundles arranged in a more parallel manner (c
and g). The bone marrow cavity was mainly occupied by adipose cells, amongst
which a great number of blood vessels were well represented. Cartilage remnant was
hardly seen inside of the scaffolds (d and h). (BM: bone marrow; Black arrow:
cartilage.)
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3.2.2.2 HA/TCP scaffold
For HA/TCP scaffolds, bone formation occurred throughout the whole implants.
Only small amount of cartilage remained after 4 weeks of in vivo implantation.
The bone tissue was characterized by close apposition of matrix on the surface
of the HA/TCP pores together with islet or bridge like bone tissue in the middle
space of the pores (Fig. 4c and d). At each in vivo time point, the length of the in
vitro priming time did not show obvious influences on the structure of the newly
formed bone and cartilage remnant.
The maturity of the bone tissue was observed to change over time in vivo. After 4
weeks of in vivo implantation, the new bone displayed in an irregular shape and the
osteoblasts could be recognized aligning at the periphery of the bone matrix (Fig.
6a and e). The cartilage remnants were observed to fill small pores or mix with the
bone tissue (Fig. 6b and f). Marrow stroma development was initiated and filled
with abundant hematopoietic cells. After 8 weeks in vivo, the bone tissue further
matured. As shown in Fig. 6c and g, the matrix exhibited a concentric pattern and
the collagen bundles arranged in a more parallel manner compared to the early
stage bone formation. Bone surface was partially covered by elongated, thin lining
cells. Moreover, the bone marrow cavity was mainly occupied by adipose cells,
amongst which a great number of blood vessels were well represented (Fig. 6d
and h). Cartilage remnants were hardly detected inside of the scaffolds.
3.2.3 Histomorphometrical evaluation
The EVG staining and Collagen II staining were used to visualize the presence of
bone and cartilage tissue, respectively. As shown in the overview pictures of PLGA/
PCL scaffolds (Fig. 7a), longer in vitro priming time resulted in more homogeneous
distribution of in vivo bone formation and less fibrous tissue occupation within the
PLGA/PCL scaffolds. The collagen II staining confirmed the significant existence of
the remaining cartilage matrix (Fig. 7b). For HA/TCP scaffolds, similar amount and
location of bone formation was revealed at all in vivo time points with very limited
presence of cartilage remnant (Fig. 8).
Quantitative results of bone and cartilage amount are depicted in Fig. 9. The
volume of newly formed bone in PLGA/PCL scaffolds did not alter significantly
within the in vivo implantation time, with the average bone area fraction ranged
from 12% to 18% (Fig. 9a). In contrast, for HA/TCP scaffolds, bone formation was
increased as the in vivo time and the average area fraction value raised from 15%
to 25% (Fig. 9b). For both scaffolds, the length of in vitro priming time did not show
obvious influences on bone volume at each in vivo time point. However, significant
higher amount of cartilage remnants was detected in PLGA/PCL scaffolds when
the constructs received 4-week chondrogenic treatment in vitro (Fig. 9c). In HA/
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Effects of In Vitro Priming Time on Endochondral Bone Formation
TCP scaffolds, marginal amount of cartilage was found after 4 weeks of in vivo
implantation and maintained stable thereafter, of which the area fraction was less
than 5% (Fig. 9b).
6
Figure 7. Overview of in vivo bone formation and cartilage remnant in PLGA/PCL
constructs. The decalcified in vivo samples were stained with EVG (a) and collagen II
staining (b) to visualize the presence of bone and cartilage tissue, respectively. Longer
in vitro priming time resulted in more homogeneous distribution of in vivo bone
formation (stained red) and less fibrous tissue occupation within the PLGA/PCL
scaffolds (a). The collagen II staining confirmed the significant existence of the
remaining cartilage matrix (stained red) (b). (scale bar: 500µm)
4. Discussion
To implement endochondral bone formation route for the clinical applications,
it is imperative to utilize an essentially shortened in vitro priming procedure
for the creation of a cartilage template to induce bone formation in vivo. In the
current study, we aimed to investigate the influence of chondrogenic priming time
on the in vivo endochondral bone formation regarding its quality and quantity.
Furthermore, to examine if such influence is universal or also depends on which
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Figure 8. Overview of in vivo bone formation and cartilage remnant in HA/TCP
constructs with different in vitro priming time. The decalcified in vivo samples were
stained with EVG (a) and collagen II staining (b) to visualize the presence of bone and
cartilage tissue, respectively. Different in vitro priming time resulted in similar
amounts and locations of bone formation (stained red) with marginal cartilage
remnants at each in vivo time point (stained red). (scale bar: 500µm)
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Effects of In Vitro Priming Time on Endochondral Bone Formation
scaffold material is applied, two scaffolds with distinguished characteristics were
selected as cell carrier, i.e. fibrous PLGA/PCL and porous HA/TCP scaffolds.
By varying the chondrogenic priming time, we were able to establish cartilage
constructs with distinct features for this study. With longer priming time, the
amount and homogeneity of cartilage matrix deposition increased significantly
over time, and meanwhile the number of undifferentiated MSCs decreased.
The expression of Col10α1, associated with chondrocyte hypertrophy [17, 18], was
detected after 2 weeks of chondrogenic priming for both PLGA/PCL and HA/TCP
constructs. These data suggest that a part of the MSCs on these scaffolds may have
proceeded toward the hypertrophic stage at this time point. For 4-week primed
constructs, a significant decrease of Col10α1 expression was shown, indicating the
late stage of cell hypertrophy. Notably, for both scaffolds, Vegfa expression of the
seeded cells has already been detected after 2 weeks, and the level maintained
stable thereafter. These data confirmed the capability of VEGF release for all
implanted constructs irrespective of the in vitro priming time.
After implantation, our data suggested that further differentiation of the MSCscaffold based cartilage constructs prior to implantation did not result in a faster or
enhanced bone formation in vivo. Similar quantity and quality of new bone tissue
was generated in vivo in HA/TCP scaffolds irrespective of in vitro culturing time.
For PLGA/PCL scaffolds, although longer-primed cartilage templates gave rise to
higher quantity of cartilage remains and more homogeneously distribution of the
new bone formation in vivo, the bone amount and maturity was in fact comparable.
These results are inconsistent with several previous studies. Scotti et al. applied an
advanced in vitro chondrogenic priming of human MSC pellets for 5 weeks, which
resulted in accelerated in vivo bone formation compared to short priming for 2
weeks [13]. In comparison, Freeman et al. chondrogenically differentiated human
and murine MSCs in pellets with different time periods from 10 days to 4 weeks.
The results demonstrated 2 or 3-week priming was optimal to achieve the highest
in vitro mineralization capacity of the cell pellets rather than 4 weeks [14]. However,
it has to be noted that, in these studies the in vivo endochondral bone formation
was tested via a scaffold-free cell pellet form, where the microenvironment and
cellular behaviors are different from the case when a scaffold is involved.
For the first time, our results indicated that, by utilizing a chondrogenically
differentiated MSC-scaffold construct as cartilage template, the essential in vitro
priming time could be shortened to 2 weeks to generate a substantial amount of
vascularized endochondral bone in vivo. This can be closely correlated to the profile
of Vegfa expression, which has already been detected after 2 weeks. As previously
shown by Farrell et al., a considerable amount of VEGF was released from the cell
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pellets after 2 weeks of in vitro chondrogenic induction [4]. After implantation, the
distinctive roles of VEGF could be exerted in the cartilage remodeling and bone
forming process, such as promoting angiogenesis [19], chemoattractant effects
on MSCs [20] and mediating osteoblastic differentiation [21]. Moreover, it is likely
that the undifferentiated MSCs also participated in the process of endochondral
bone formation. An explanation is that, these cells committed to chondrogenic
Figure 9. Quantification of bone and cartilage amount after implantation. The total
ratio of newly formed bone in PLGA/PCL scaffolds remained constant within the in
vivo implantation time (a), while for HA/TCP scaffolds, bone formation was increased
over time (b). For both scaffolds, the length of in vitro priming time did not show
obvious influences on bone volume at each in vivo time point (a and b). However,
significant higher amount of cartilage remnant was detected in PLGA/PCL scaffolds
when the constructs received 4-week chondrogenic treatment in vitro (c). In HA/TCP
scaffolds, marginal amount of cartilage was detected (d). (* p<0.05, ** p<0.01; #
represents significant difference compared to the bone ratio at week 4 in vivo, #
p<0.05, ## p<0.01; error bars represent standard deviation of the mean, n=3)
lineage after implantation and fulfill their role in hypertrophy together with
other chondrocytes to facilitate the vascularization and the recruitment of
osteoprogenitors from the host animal [2, 22]. Another explanation is, that the
undifferentiated MSCs were instructed by the hypertrophic chondrocytes to
become osteoblasts or hematopoietic cells, and directly participated in building
the vascularized bone tissue [23]. In either case, it implies that a certain extent
of in vitro chondrogenic priming of the MSCs, e.g. for 2 weeks, or possibly with
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Effects of In Vitro Priming Time on Endochondral Bone Formation
even shorter time, might be sufficient to ensure the progression of endochondral
differentiation of the seeded cells and enable successful bone formation, even
after been implanted in a non-chondrogenic or non-osteogenic environment.
Instead of being influenced by the in vitro priming time, our data have suggested
that the newly formed endochondral bone could be generated in different manners
based on the chemical and structural cues provided by the scaffold design. Newly
formed woven bone like structure and significant cartilage presence have been
observed in PLGA/PCL scaffolds after 4 weeks in vivo. Surprisingly, such structures
were maintained up to 8 weeks. In comparison, cartilage remnants were hardly
seen in HA/TCP scaffolds after 4 weeks. Over time, the amount and maturity of
endochondral bone formation significantly improved after 8 weeks. These findings
indicated that, HA/TCP scaffold supported more rapidly progression of the
endochondral bone generation and vascularization compared to PLGA/PCL, which
might be closely related to the high VEGF expression of the cells stimulated by
the HA/TCP scaffolds. Regarding the pattern of the bone matrix, HA/TCP scaffolds
instructed the new bone tissue to deposit on the surface of the interconnected
pores in a polarized fashion, whereas for PLGA/PCL constructs, a fully integrated
bone/cartilage/electrospun fiber complex was built without any preferential
arrangement of bone tissue. Similar results have also shown previously by Scaglione
et al. and they addressed such different bone-forming patterns were attributed to
the specific scaffold geometry [24]. The confined porous structure could concentrate
collagen fibers and serve as a polarizing template within the pores to trigger the
formation of lamellar like bone. On the contrary, fibrous scaffolds possessed a
high level of geometrical complexity, thus the new collagen fibers formed by the
cells were distributed in a isotropic manner, resulting mostly in woven bone [24].
Above all, our data confirmed that scaffolds played a crucial role in influencing
cell behaviors and bone deposition kinetics. The control over endochondral bone
formation could be exerted through rational design of the scaffolds.
This study contributes to the view that chondrogenical priming of rat MSCs in the
presence of biomaterials followed by in vivo implantation was able to generate
endochondral bone in vivo. Moreover, we demonstrated that it was not necessary
to differentiate MSCs till fully hypertrophy to accomplish vascularized bone
formation. However, it has to be noted that this conclusion needs to be further
evidenced in a more clinically relevant situation by using human MSCs and
implanting in an orthotopic location, where the mechanical load and biochemical/
inflammatory factors are closely involved. Moreover, we could not conclude if
2-week is the essential time length for chondrogenic priming and if such process
can be further shortened, because no other shorter period has been chosen for
this study. Additionally of interest is the further understanding of 1) the function
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and the fate of the implanted cells, and 2) the specific responses of the seeded
cells induced by particular scaffold properties and the underlying mechanisms, so
that the endochondral route can be established to target efficient and large-scale
bone regeneration.
5. Conclusion
In vitro chondrogenic priming of rat MSCs for 2, 3 and 4 weeks leads to vascularized
endochondral bone formation in vivo on two distinct scaffolds, i.e. fibrous PLGA/
PCL and porous HA/TCP scaffolds. A longer in vitro priming time can result in a
more homogeneous distribution of bone formation, whereas this does not
have to give rise to a higher amount of in vivo bone formation in the selected
scaffolds. Evidently, in our study the quality of endochondral bone was influenced
by the chemical-physical design of the scaffolds. An unpolarized woven bone
like structure was generated in fibrous PLGA/PCL scaffolds, while porous HA/TCP
scaffolds supported progressive lamellar-like bone formation with mature bone
marrow development. Further investigations are warranted to investigate the
essential in vitro priming time and the mechanism of cell-scaffold interaction in
the endochondral bone forming process.
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Chapter 7
Roles of cartilage templates in
mediating mesenchymal stem
cell behavior in endochondral
bone tissue engineering
Wanxun Yang, Sanne K. Both, John A. Jansen and Fang Yang
Roles of Cartilage Templates in Mediating MSCs Behavior
Summary of this chapter
Costly and lengthy process for creating cartilage template hampers the translation
of endochondral approach to clinic. A more comprehensive understanding of
the crosstalk between implanted cartilage template and host cells is essentially
beneficial to possibly simplify this procedure or develop substitutes for effective
bone regeneration. Therefore, this chapter aimed to explore the role of cartilage
template in mediating MSC behavior in the process of endochondral bone
formation, as well as the possible signaling pathways involved in these biological
events. To that end, rat bone marrow derived mesenchymal stem cells (MSCs) were
seeded on a porous hydroxyapatite (HA) /tricalcium phosphate (TCP) composite
and cultured in chondrogenic condition to create the cartilage templates.
Afterwards, undifferentiated MSCs were co-cultured with these constructs in
a transwell system (separating cartilage template and MSCs with a permeable
membrane) to assess the migration and osteogenic differentiation capacity of
the MSCs. Furthermore, the expression profile of chemokine signaling related
genes was analyzed via PCR array. The results demonstrated that chondrogenically
differentiated MSC-HA/TCP constructs (chondro-TCP) were able to induce MSCs
migration by possibly activating the IL8/CXCR2 signaling pathway. Additionally,
they could trigger a considerable elevation of signaling factors responsible for
MSCs recruitment and osteoclast stimulation, including CCL3, CCL7, CCL19 and NFκB. The osteogenic differentiation of MSCs was not significantly influenced after
short exposure to chondro-TCP templates.
1. Introduction
During both embryogenesis and postnatal fracture repair, the majority of bone
formation is governed by a mechanism known as endochondral ossification [1].
Such endochondral process has recently inspired a new concept of bone tissue
engineering via a two-step approach, during which a cartilage intermediate is first
created in vitro, and subsequently implanted as an osteoinductive template in
vivo to transform into bone [2]. It has been revealed this endochondral approach
has a superior efficacy for generating vascularized bone tissue compared to
the traditional intramembranous ossification method [3-5]. Still, several issues
hurdle the effective translation of this tissue engineering strategy to the clinics.
For instance, the creation of cartilage templates is usually time-consuming (3-5
weeks), and requires a large cell number as well as high usage of growth factors,
e.g, transforming growth factor beta (TGF-β) and bone morphogenetic proteins
(BMPs). All of these associate not only the difficulties in obtaining sufficient
cell source but also considerably high treatment costs and regulatory burdens.
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Therefore, it is necessary to simplify the in vitro procedure or develop alternative
products/compounds with similar or better functions as cartilage templates.
Essentially, this would not be possible without a complete understanding of
the mechanism underlying this endochondral approach, especially the in vivo
functions of the cartilage template.
In practice, the engineered cartilage template used in endochondral approach is
usually created by loading mesenchymal stem cells (MSCs) in a three-dimensional
scaffold or in cell pellet form followed by in vitro chondrogenic culturing. After
implantation, the transformation from cartilage to bone is a very complex
progression and involves the coordination of a variety of cell populations. Previous
studies have shown that, instead of the direct contribution of the seeded cells
from the cartilage templates, the in vivo endochondral bone formation was
predominately originated by host cells [6]. Based on these results, it is speculated
that the cartilage template plays an essential role in recruiting the host cells and
regulating their performance in the successive bone forming process. Yet, this
hypothesis has not been tested. The exact functions of the cartilage template
for inducing regenerative functions of the host cells, e.g. MSCs, and the involved
signaling pathways still remain unknown.
Therefore, this study intended to investigate the effects of cartilage template
on MSCs in terms of their migration and osteogenic differentiation ability.
Furthermore, we also aimed to discover the possible signaling pathways involved
in the crosstalk between cartilage templates and host MSCs. It is known that the
communications between the cell populations are realized by different ways,
including cell-cell, cell-matrix interactions, release of the soluble factors, or a
combination thereof. Our study represented the first step and mainly focused on
the effects of soluble factors. To that end, rat bone marrow derived MSCs were
first seeded on a porous hydroxyapatite/tricalcium phosphate composite (HA/TCP)
and cultured in chondrogenic condition to create the cartilage templates based on
our previous study [5]. Afterwards, undifferentiated MSCs were co-cultured with
these constructs in an indirect co-culture system (separating cartilage template
and MSCs with a permeable membrane) to assess the migration and osteogenic
differentiation capacity of the MSCs. Furthermore, the expression profile of
chemokine signaling related genes was analyzed via PCR array to investigate the
mechanism underlying the biological response of MSCs elicited by the cartilage
template.
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Roles of Cartilage Templates in Mediating MSCs Behavior
2. Materials and methods
2.1 Cell isolation
Four 6-week-old male Wistar rats (Charles River, Wilmington, USA.) were used
for MSC isolation after the approval from Radboud University Animal Ethics
Committee. Two femora of each rat were removed and the epiphyses were cut
off. Afterwards, rat MSCs were flushed out of the remaining diaphyses using
the primary culture medium through an 18G needle (BD Microlance, Drogheda,
Ireland), which consisted of alpha Minimal Essential Medium (αMEM; Gibco,
Grand Island, USA) supplemented with 10% fetal bovine serum (FBS; Sigma, St.
Louis, USA), 100 U/ml penicillin and 100μg/ml streptomycin (Gibco). The flushout was cultured for 2 days in a humidified incubator (37 °C, 5% CO2), after which
the medium was refreshed to remove non-adherent cells. The cells were cultured
for an additional 3 days followed by detachment using trypsin/EDTA (0.25% w/v
trypsin, 0.02% w/v EDTA; Sigma) and counted. Half of the obtained the cells were
used for cell seeding on the scaffolds, the other half was frozen in the culture
medium containing 10% DMSO in the liquid nitrogen condition for later use.
2.2 Preparation of cartilage template
A porous HA/TCP composite (CAM bioceramics B.V., Leiden, the Netherlands)
was used as cell carrier, which was obtained in a cubic shape of 3×3×3 mm3
and consisted of HA and TCP with a 75/25 ratio. The volumetric porosity was
75% (provided by the manufacturer). Prior to cell seeding, the scaffolds were
autoclaved and soaked in the primary culture medium overnight. During cell
seeding, the scaffolds were incubated in the cell suspension with a concentration
of 1 × 106 cells/ml (4 scaffolds per 1 ml of cell suspension) and gently rotated
for 3 hours. Afterwards, the scaffolds were placed in non-adherent tissue culture
plates and the unattached cells from the suspension were collected, centrifuged
and reseeded onto the scaffolds to ensure a high cell loading efficiency. All of the
constructs were proliferated for 1 week. Thereafter, the medium was changed to
chondrogenic medium, consisting of high-glucose DMEM (Gibco), 1% FBS, 100 U/
ml penicillin, 100 μg/ml streptomycin, 50 μg/ml l-ascorbic acid, 100 mM of sodium
pyruvate (Gibco), 1:100 insulin-transferrin-selenium (ITS; BD Biosciences, Bedford,
USA), 100nM dexamethasone, 10 ng/ml of transforming growth factor beta-2
(TGF-β2; R&D systems, Minneapolis, USA) and 100 ng/ml bone morphogenetic
protein 2 (BMP-2; BD biosciences). The constructs were cultured for 4 weeks
to obtain the cartilage templates (named chondro-TCP). Bare HA/TCP scaffolds
without cell seeding were used as control (named bare-TCP).
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2.3 MSCs migration test
A tranwell migration model was used to examine the migration of MSCs in response
to the cartilage constructs (Fig. 1a). Primary rat MSCs were defrosted, expanded
and then seeded onto ThinCertTM cell culture inserts (membrane pore size 8μm,
Greiner Bio-One) in a 24-well plate at a density of 5,000 cells/insert. Then 200 μl
and 600 μl DMEM supplemented with 1% FBS was added into the insert and lower
chamber, respectively. Subsequently, two chondro-TCP or bare-TCP constructs
were placed in the lower chambers. After 24 h culture in a humidified incubator
(37 °C, 5% CO2), the inserts were removed and the upper side of the membrane
was scraped to remove adherent cells. Further, they were fixed in 10% formalin for
30 min, and stained with toluidine blue (0.05%) for 30 min to visualize the migrated
cells on the lower side of the membrane. All samples were then photographed
with Zeiss Imager Z1 equiped with AxioCam MRc5 camera (Carl Zeiss Microimaging
GmbH, Göttingen, Germany). The number of the migrated cells was quantified
based on the obtained images using Image J software (National Institutes of
Health, Bethesda, USA).
Figure 1. Scheme of co-culture systems.
2.4 Osteogenic differentiation of MSCs
To further investigate the influence of cartilage constructs on the osteogenic
differentiation of MSCs, a second indirect co-culture system with smaller
membrane pore size was established (Fig. 1b). MSCs were seeded in the 24-well
plate at a density of 100,000 cells per well. Afterwards, two chondro-TCP or bareTCP constructs were added into the top inserts (ThinCertTM cell culture inserts,
membrane pore size 0.4 μm, Greiner Bio-One) overlying the MSCs. Two-hundred
μl and 800 μl DMEM supplemented with 1% FBS was added into the insert and the
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Roles of Cartilage Templates in Mediating MSCs Behavior
lower well, respectively. After 4 days, MSCs in the lower well were collected for
the following tests.
2.4.1 PCR analysis of osteogenic-related genes
RNA of MSCs was extracted using TRIzol® reagent according to the manufacturer’s
instructions (Life Technologies, Carlsbad, USA). The obtained RNA (n=3) was
dissolved in RNase free water and the concentration was measured with
spectrophotometer (NanoDrop 2000, Thermal scientific, Wilmington, USA).
Subsequently, first strand cDNA was reverse transcribed from RNA using iScript™
cDNA Synthesis Kit (Bio-Rad Laboratories, Inc., USA). Afterwards cDNA was further
amplified and the expression of specific genes was quantified using IQ SYBR Green
Supermix PCR kit (BioRad, Hemel Hempstead, United Kingdom) in a real-time PCR
(BioRad, CFX96™ real-time system) (n=4). Osteogenic-related gene markers were
evaluated, including runt-related transcription factor 2 (Runx2), collagen type 1
(Col1α1), alkaline phosphatase (Alp), and bone sialoprotein (Bsp) (Table 1). The
specificity of the primers was confirmed separately before the real-time PCR
reaction. The expression levels were compared to the housekeeping gene β-Actin
and calculated via the 2-∆∆Ct method [7].
Table 1. Overview of the primer sequences
Forward (5’ → 3’)
Reverse (5’ → 3’)
Runx-2
GAGCACAAACATGGCTGAGA
TGGAGATGTTGCTCTGTTCG
Col1α1
GAGCGATTACTACTGGATTGACCC
CAAGGAATGGCAGGCGAGAT
Alp
GGGACTGGTACTCGGATAACGA
CTGATATGCGATGTCCTTGCA
Bsp
TCCTCCTCTGAAACGGTTTCC
GGAACTATCGCCGTCTCCATT
β-Actin
TTCAACACCCCAGCCATGT
TGTGGTACGACCAGAGGCATAC
7
2.4.2 DNA content and ALP activity
DNA content (n=3) of MSCS was quantified by QuantiFluorTM DNA system (Promega
Corporation, Madison, USA) according to the manufacturer’s instructions. Culture
medium and upper inserts was removed, and cells were washed twice with PBS.
One ml of milliQ water was added to each well, after which 2 freeze-thaw cycles
were performed. A DNA standard curve was made from the Picogreen assay kit
(Quant-iT PicoGreen dsDNA, Invitrogen). 100 μl sample or standard were added
to 100 μl freshly made working solution in the wells. The results were read using a
fluorescence microplate reader (Bio-Tek Instruments, Abcoude, The Netherlands)
with an excitation wavelength at 485nm and an emission wavelength at 530nm.
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Chapter 7
The ALP activity (n=3) was measured as a marker for osteogenic differentiation
using the same cell extracts as for the DNA assay. 80 μl of sample or standard
and 20μl buffer solution (5 mM MgCl2, 0.5 M 2-amino-2-methyl-1-propanol)
was pipetted into a 96-well plate, and 100μl of substrate solution (5 mM
paranitrophenylphosphate) was added per well. Subsequently, the plate was
incubated for 1 h at 37 °C, after which the reaction was stopped by adding 100 μl of
0.3M NaOH. Serial dilutions of 4-nitrophenol (0–250 ng/ml) were used for making
the standard curve. The plate was read in an ELISA reader (Bio-Tek Instruments) at
405 nm. ALP activity results were normalized by the amount of DNA.
2.5 PCR array of chemokine signaling related genes
The expression profile of chemokine signaling related genes of MSCs was analyzed
by pathway based quantitative PCR arrays (RT2Profiler PCR arrays, Rat Chemokines
& Receptors, no. PAMM-022, SABiosciences) using the same RNA extration as
described in 2.4.1. RT2 real-time SYBR Green PCR master mix (SABiosciences) was
applied according to the manufacturer’s instructions. The tested genes for the PCR
array were listed in Supplementary Table 1. The protocol used for PCR array was
95°C for 10 min, followed by 40 cycles of 95 °C for 15 sec and 60°C for 1 min. The
data were analyzed using the SABiosciences PCR Array Data Analysis Software.
Figure 2. Cell migration. Compared to bare-TCP scaffolds (c), chondro-TCP constructs
(b) significantly promoted MSC migration, which resulted in a 4-fold change (p<0.05)
of cell number across the transwell membrane (a). (staining: toluidine blue; error
bars represent standard deviation of the mean, n=3)
3. Results
3.1 MSCs migration
In vitro rat MSCs recruitment was tested using a transwell system. Toluidine blue
staining showed that chondrogenically differentiated HA/TCP-cell constructs were
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Roles of Cartilage Templates in Mediating MSCs Behavior
able to induce MSCs migration across the transwell membrane (Fig. 2). Compared
to the bare-TCP scaffolds, chondro-TCP significantly enhanced rat MSCs migration
(p<0.05) by 4-fold.
3.2 Expression of osteogenic genes
The expression of osteogenic gene markers of MSCs was detected after 4 days
of co-culture. As shown in Fig. 3, for all detected genes, including Runx2, Col1a1,
Alp and Bsp, there was no significant difference between MSCs co-cultured with
chondro-TCP and MSCs with bare-TCP constructs.
Figure 3. Osteogenic gene expressions of MSCs after 4 days of co-culture.
7
Figure 4. ALP activity (a) and DNA content (b) of MSCs after 4 days of co-culture.
139
Chapter 7
3.3 DNA content and ALP activity
After 4 days of co-culture, ALP activity could be detected from MSCs which were
co-cultured with chondo-TCP or bare-TCP constructs, and the amounts were
comparable. Similar DNA content, which is related to the cell number, was also
shown for these two groups (Fig. 4).
3.4 Analysis of chemokine signaling pathway
PCR array was used to identify the genes that might be involved in the behavior
change of MSCs in response to the cartilage template. The result revealed that
4-day co-culture of MSCs with chondro-TCP resulted in the expression level
alteration of a number of genes for MSCs (Fig. 5 and supplementary Table. 1),
among which several chemokines, i.e. Ccl3, Ccl7, Ccl19 and Nfkb1, and receptors,
i.e. Cxcr2 and Tlr2, were considerably up-regulated (gene fold change > 4) (Table.
2).
Figure 5. PCR array of chemokine signaling pathway. Each square represents one
tested gene and its color indicates the gene fold change (up- or down-regulation) of
MSCs co-cultured with chondro-TCP compared to MSCs co-cultured with bare-TCP.
4. Discussion
Although endochondral ossification approach showed significant superiority
in generating vascularized bone for tissue engineering applications, costly and
lengthy process for creating cartilage template hampers its translation to clinic. A
more comprehensive understanding of the crosstalk between implanted cartilage
140
Roles of Cartilage Templates in Mediating MSCs Behavior
template and host cells is essentially beneficial to possibly simplify this procedure
or develop substitutes for effective bone regeneration. Therefore, the aim of the
current study was to analyze the effect of cartilage template on bone marrow
derived MSCs and also the possible signaling pathways involved in these biological
events.
Table 2. Overview of up- and down-regulation of the genes in PCR array (Numbers
indicate the gene fold change compared to the control group).
Up-regulation (gene fold change > 2)
Down-regulation (gene fold change > 2)
Ccl19 (14.5), Tlr2 (10.1), Cxcr2 (8.5), Ccl3 (8.1), Ccl7
Bmp10 (5.0), Il13 (2.11)
(7.1), Nfkb1 (4.8), Ccbp2 (3.5), Cmklr1 (2.8), Ccl6 (2.8),
Tnf (2.7), Csf2 (2.5), Bdnf (2.3), Cxcl13 (2.3), C5 (2.2),
Gdf5 (2.1)
Our results showed that cartilage templates (chondro-TCP constructs) were
able to induce MSCs migration when tested using a transwell culture system,
which suggested their possible role for MSC recruitment in vivo. It is known that
interactions of chemokines and their receptors mediate the migration of MSCs.
Among various chemokines, stromal cell-derived factor-1 (SDF-1), also known as
CXCL12, has been reported to play a pivotal role in MSCs homing and modulating
[8, 9]
. SDF-1 activates cell recruitment mainly through its receptor CXCR4 [10]. The
interaction between SDF-1 and CXCR4 also stimulates the release of matrix
metalloproteinases (MMPs) [11, 12], which are closely involved in the remodeling of
cartilage matrix [13]. However, in our study, the presence of chondro-TCP constructs
did not activate higher CXCR4 expression of the tested MSCs compared to the bareTCP control. Instead, a significant up-regulation was found for CXCR2 expression,
the receptor of interleukin 8 (IL8), indicating that the observed MSCs migration
might be associated with IL8 signaling. This result is supported by some previous
investigations, which also demonstrated the effect of IL8 on MSCs homing [8, 14].
Overall, PCR array results demonstrated the up-regulation of a number of
chemokine signaling related genes, indicating MSCs were actively stimulated
by the cartilage templates. Amongst the significantly up-regulated genes, CCL7
(also known as monocyte-specific chemokine 3; MCP-3), CCL 19 (also known as
macrophage inflammatory protein-3β; MIP-3β), CCL3 (also known as macrophage
inflammatory protein-1α; MIP-1α), and NF-κB (nuclear factor kappa B) were of
particular interest. The involvement of CCL 3, CCL7 and CCL19 in migration of
MSCs and hematopoietic stem cells (HSCs) have been reported previously [15, 16].
Up-regulation of these three genes suggested that, besides being recruited by
141
7
Chapter 7
the cartilage templates, MSCs were also stimulated to express more inducers to
recruit additional cells participating in the successive actions. In addition to cell
recruitment, the up-regulated genes (CCL3, CCL19 and NF-κB) are also closely
related to the regulation of osteoclast functions [17-19]. In particular, the activation
of NF-κB pathway, which is commonly mediated by RANKL (receptor activator of
the NF- κB ligand), might significantly promote osteoclastogenesis [19]. As a result,
the remodeling of cartilage matrix could be directly enabled. These data suggested
that, short-term release of soluble mediators from cartilage templates were
able to effectively trigger a cascade of genes activations of MSCs, which would
potentially contribute to the bone forming process by recruiting additional cells
and regulating osteoclast activities.
Regarding the osteogenic differentiation of MSCs, the presence of chondroTCP constructs did not show any significant effects in a short term. This was
evidenced by the PCR results that comparable osteogenic gene expressions and
ALP activities were detected for both chondro-TCP and bare-TCP group after 4
days of co-culture. Despite all that, the cartilage templates might influence the
osteogenic differentiation of MSCs at a later time point. As demonstrated in some
previous studies, chondrocytes were able to significantly induce the osteogenic
differentiation of MSCs by continuous co-culture of 21 days [20, 21]. Another
possibility is, the released soluble factors alone might not be sufficient to support
osteo-induction of MSCs. A stronger osteogenic induction for MSCs would possibly
necessitate soluble factors in conjunction with a direct cartilage-MSCs contact [22].
Some technical remarks should be made regarding the study set-up. Although coculture systems have been commonly explored in vitro to investigate the interactions
between different cell populations in a manner of maximally simulating the in vivo
complexity, it is still very difficult to replicate the real situation. One of the main
concerns is the choice of co-culture medium. Varied culture conditions, i.e. fullor half-supplemented culture medium (chondrogenic or osteogenic), have been
applied previously in co-cultures studying the interaction between chondrocytes
and MSCs [21, 23]. However, this would inevitably bring too many variables, and also
have the risk in eliciting artificially high response of MSCs from the supplements.
Therefore, in our study, the co-culture was performed in a plain medium. In this
way, the low-serum culture condition for cartilage template could be maintained,
and the effects of cartilage templates on MSCs could be evaluated without the
possible disturbances from the osteogenic supplements (i.e. ascorbic acid, Naβ-glycerolphosphate or dexamethasone). Nevertheless, it has also to be noted
that such low nutrition culture condition, in some extent, might also influence
the survival and function of both MSCs and cartilage templates. Therefore, the
142
Roles of Cartilage Templates in Mediating MSCs Behavior
exploration of a more suitable co-culture condition (or system) is required for the
future studies.
Besides the soluble mediators, follow-up studies are warranted to investigate
the influence of cartilage templates on MSCs in form of direct contact. Another
limitation of our study is, the co-culture was only performed in a short period
due to the consideration of cell survival in the plain culture medium. Therefore,
long-term effect of cartilage constructs on MSCs need to be further investigated.
Moreover, as the in vivo endochondral bone forming process involves several other
cell types in addition to MSCs and chondrocytes, e.g. a variety of hematopoietic
cells, a more comprehensive understanding about the interaction between
cartilage template and other cell types is also necessary. When the crosstalk
between implanted cartilage and host cells is thoroughly unveiled, endochondral
bone engineering can be further simplified and more effectively applied, i.e. by
using scaffolds loaded with biological cues that provoke host cell performance
without implantation of exogenous cells. The data of this study represent the first
step in this process.
5. Conclusion
By using an indirect co-culture system, our results demonstrated that
chondrogenically differentiated MSC-HA/TCP constructs (chondro-TCP) were able
to induce MSCs migration by possibly activating the IL8/CXCR2 signaling pathway.
Additionally, they could activate a considerable elevation of signaling factors
responsible for MSCs recruitment and osteoclast stimulation, including CCL3, CCL7,
CCL19 and NF-κB. The osteogenic differentiation of MSCs was not significantly
influenced after short exposure to chondro-TCP templates. This study represents
the first step determining the possible signaling pathways through which the
cartilage templates exert their functions in generating endochondral bone tissue.
Continued research in this direction would open the attractive possibility for the
development of alternative “cartilage templates” to target endochondral bone
regeneration.
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Roles of Cartilage Templates in Mediating MSCs Behavior
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7
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Chapter 7
Supplementary Table 1: Gene table of PCR array (rat chemokines & receptors)
Position
Gene
Description
symbol
Gene
Fold up- or
reference
down-
NM_031349
NM_012513
NM_053303
NM_001031824
NM_021670
NM_013107
NM_016994
XM_342421
NM_078621
NM_019205
NM_001105822
NM_057151
NM_001108661
NM_031530
NM_019233
NM_013025
NM_053858
NM_031116
NM_001004202
NM_001007612
NM_001012357
NM_020542
NM_001106872
NM_021866
NM_053958
NM_133532
NM_053960
NM_001013145
NM_199489
XM_236704
NM_172329
NM_001108191
NM_001013142
NM_001106034
NM_022218
NM_053352
NM_001106164
NM_023981
XM_340799
NM_134455
NM_133534
NM_030845
NM_139089
NM_182952
NM_053647
NM_001007729
NM_022214
NM_153721
2
2.29
1.85
-5.01
2
-1.62
1.69
2.19
3.52
1.2
2
2
14.48
2
1.38
8.13
1.12
1.7
2.8
7.14
2
-1.15
-1.06
2
1.01
2
1.1
2
2
2
2
2
-1.18
1.99
2.75
1.37
1.28
1.8
2.5
-1.27
1
2
1.13
1.28
-1.42
1.19
2
1.41
regulation
A01
A02
A03
A04
A05
A06
A07
A08
A09
A10
A11
A12
B01
B02
B03
B04
B05
B06
B07
B08
B09
B10
B11
B12
C01
C02
C03
C04
C05
C06
C07
C08
C09
C10
C11
C12
D01
D02
D03
D04
D05
D06
D07
D08
D09
D10
D11
D12
146
Aplnr
Bdnf
Cxcr5
Bmp10
Bmp15
Bmp6
C3
C5
Ccbp2
Ccl11
Ccl12
Ccl17
Ccl19
Ccl2
Ccl20
Ccl3
Ccl4
Ccl5
Ccl6
Ccl7
Ccl9
Ccr1
Ccr1l1
Ccr2
Ccr3
Ccr4
Ccr5
Ccr6
Ccr7
Ccr8
Ccr9
Ccrl2
Cmtm2a
Cmtm5
Cmklr1
Cxcr7
Cmtm3
Csf1
Csf2
Cx3cl1
Cx3cr1
Cxcl1
Cxcl10
Cxcl11
Cxcl2
Pf4
Cxcl5
Ppbp
Apelin receptor
Brain-derived neurotrophic factor
Chemokine (C-X-C motif) receptor 5
Bone morphogenetic protein 10
Bone morphogenetic protein 15
Bone morphogenetic protein 6
Complement component 3
Complement component 5
Chemokine binding protein 2
Chemokine (C-C motif) ligand 11
Chemokine (C-C motif) ligand 12
Chemokine (C-C motif) ligand 17
Chemokine (C-C motif) ligand 19
Chemokine (C-C motif) ligand 2
Chemokine (C-C motif) ligand 20
Chemokine (C-C motif) ligand 3
Chemokine (C-C motif) ligand 4
Chemokine (C-C motif) ligand 5
Chemokine (C-C motif) ligand 6
Chemokine (C-C motif) ligand 7
Chemokine (C-C motif) ligand 9
Chemokine (C-C motif) receptor 1
Chemokine (C-C motif) receptor 1-like 1
Chemokine (C-C motif) receptor 2
Chemokine (C-C motif) receptor 3
Chemokine (C-C motif) receptor 4
Chemokine (C-C motif) receptor 5
Chemokine (C-C motif) receptor 6
Chemokine (C-C motif) receptor 7
Chemokine (C-C motif) receptor 8
Chemokine (C-C motif) receptor 9
Chemokine (C-C motif) receptor-like 2
CKLF-like MARVEL transmembrane domain containing 2A
CKLF-like MARVEL transmembrane domain containing 5
Chemokine-like receptor 1
Chemokine (C-X-C motif) receptor 7
CKLF-like MARVEL transmembrane domain containing 3
Colony stimulating factor 1 (macrophage)
Colony stimulating factor 2 (granulocyte-macrophage)
Chemokine (C-X3-C motif) ligand 1
Chemokine (C-X3-C motif) receptor 1
Chemokine (C-X-C motif) ligand 1
Chemokine (C-X-C motif) ligand 10
Chemokine (C-X-C motif) ligand 11
Chemokine (C-X-C motif) ligand 2
Platelet factor 4
Chemokine (C-X-C motif) ligand 5
Pro-platelet basic protein (chemokine (C-X-C motif) ligand 7)
Roles of Cartilage Templates in Mediating MSCs Behavior
E01
E02
E03
E04
E05
E06
E07
E08
E09
E10
E11
E12
F01
F02
F03
F04
F05
F06
F07
F08
F09
F10
F11
F12
G01
Cxcl9
Cxcr3
Cxcr4
Tymp
Epo
Gdf5
Ccr10
Gpr77
Hif1a
Il13
Il16
Il18
Il1a
Il4
Il8ra
Cxcr2
Inha
Inhbb
Lif
Cxcl13
Ltb4r2
Mmp2
Mmp7
Myd88
Nfkb1
Chemokine (C-X-C motif) ligand 9
Chemokine (C-X-C motif) receptor 3
Chemokine (C-X-C motif) receptor 4
Thymidine phosphorylase
Erythropoietin
Growth differentiation factor 5
Chemokine (C-C motif) receptor 10
G protein-coupled receptor 77
Hypoxia-inducible factor 1, alpha subunit
Interleukin 13
Interleukin 16
Interleukin 18
Interleukin 1 alpha
Interleukin 4
Interleukin 8 receptor, alpha
Chemokine (C-X-C motif) receptor 2
Inhibin alpha
Inhibin beta-B
Leukemia inhibitory factor
Chemokine (C-X-C motif) ligand 13
Leukotriene B4 receptor 2
Matrix metallopeptidase 2
Matrix metallopeptidase 7
Myeloid differentiation primary response gene 88
Nuclear factor of kappa light polypeptide gene enhancer in
NM_145672
NM_053415
NM_022205
NM_001012122
NM_017001
XM_001066344
NM_001108836
NM_001003710
NM_024359
NM_053828
NM_001105749
NM_019165
NM_017019
NM_201270
NM_019310
NM_017183
NM_012590
NM_080771
NM_022196
NM_001017496
NM_053640
NM_031054
NM_012864
NM_198130
XM_342346
2
2
2
2
-1.05
2.14
1.71
1.21
1.42
-2.11
1.45
1.25
2
1.13
-1.14
8.49
2
-1.11
2
2.25
1.85
2
1.22
1.77
4.79
G02
G03
G04
G05
G06
G07
G08
G09
G10
G11
G12
Rgs3
Sdf2
Slit2
Tapbp
Tcp10b
Tlr2
Tlr4
Tnf
Tnfrsf1a
Trem1
Xcl1
B-cells
Regulator of G-protein signaling 3
Stromal cell derived factor 2
Slit homolog 2 (Drosophila)
TAP binding protein
T-complex protein 10b
Toll-like receptor 2
Toll-like receptor 4
Tumor necrosis factor (TNF superfamily, member 2)
Tumor necrosis factor receptor superfamily, member 1a
Triggering receptor expressed on myeloid cells 1
Chemokine (C motif) ligand 1
NM_019340
NM_001105803
NM_022632
NM_033098
NM_001107461
NM_198769
NM_019178
NM_012675
NM_013091
NM_001106885
NM_134361
1.53
2
1.42
-1.06
1.4
10.09
1.12
2.66
-1.78
2.46
2
7
147
Chapter 8
Summary, closing remarks
and future perspectives
Summary, Closing Remarks and Future Perspectives
1. Summary and address to the aims
Although great challenges remain, bone tissue engineering presents a promising
strategy to replace the current autografts or allografts for the treatment of
bone defects. The standard procedure of bone tissue engineering involves
loading osteoprogenitor cells onto proper designed biomaterials to induce cell
differentiation in a specific pathway, followed by transplantation of this cellscaffold complex into the defect site to regenerate bone tissue. Such a construct
aims not only to replace the injured tissue, but also to induce and direct the healing
process, and completely restore the functions of the native tissue. So far, the most
common approach to generate bone has largely focused on intramembranous
ossification (i.e. bone formation through direct osteoblastic differentiation),
however without obtaining convincing results in humans. Until recently, a new
approach, resembling endochondral ossification, has drawn attention. With
this approach, bone is formed via the remodeling of a cartilage intermediate.
Compared to osteogenically differentiated cells, chondrocytes are well adapted
to limited nutrition supply and hypoxic conditions. In addition, they can produce
considerable amounts of angiogenic factors during differentiation. These features
give the endochondral approach a great advantage in maintaining viability of
the implanted cells and generating vascularized bone tissue. Yet the available
knowledge and investigations on this endochondral approach is still in its infancy.
This thesis aimed to investigate and optimize the endochondral ossification
approach to achieve high-quality bone regeneration. Specifically, its applicability
with different scaffold materials, optimal in vitro priming time, as well as the
mechanism of the transformation from cartilage to bone, were explored in a series
of studies. Chapter 1 gives a general introduction on bone tissue engineering and
its essential elements, as well as the aims of this thesis. Each following chapter
presents a separated study. This summary addresses the research questions as
described in the first chapter in successive order.
1. What is the current state of the art of combining stem cells and biodegradable
materials for bone regeneration via the endochondral pathway?
Thanks to the improving knowledge of stem cell sources, together with profound
advances in materials sciences, the “engineering” of man-made bone grafts
appears to be increasingly feasible to target the regeneration of large bone defects.
Chapter 2 first elaborates on bone tissue and osteogenesis mechanisms, in order to
reach a fundamental understanding of the processes involved in tissue engineering
approaches for bone regeneration. Next, the preliminary investigations in utilizing
biomaterials and cell culture strategies to achieve endochondral bone formation
were highlighted. Different cells sources, e.g. chondrocytes, MSCs and ESCs have
151
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Chapter 8
been reported to posses the ability of undergoing the endochondral differentiation
pathway with or without specific stimulations. A variety of biomaterials, with the
structure and composition closely mimicking the nature cartilage or bone tissue,
are potential candidates to support the generation of osteochondral tissue,
among which natural or synthetic polymers and ceramics are the top choices.
The superiority of generating vascularized bone has been revealed by using the
endochondral strategy, albeit standardization and optimization of this approach
are essentially needed to achieve an efficient transformation from cartilage to
bone tissue. Moreover, a number of issues still need to be addressed before a
clinical trial can be started, including the appropriate time point for implantation,
function and fate of implanted cells.
2. Is the porous aliphatic polyurethane/hydroxyapatite composite a suitable
scaffold for bone tissue engineering applications?
Biomaterial scaffolds meant to function as supporting structures to osteogenic
cells play a pivotal role in bone tissue engineering. In Chapter 3, we synthesized
an aliphatic polyurethane (PU) scaffold via a foaming method using non-toxic
components. Through this procedure a uniform interconnected porous structure
was created. Furthermore, hydroxyapatite (HA) particles were introduced into this
process to increase the bioactivity of the PU matrix. To evaluate the biological
performances of these PU based scaffolds, their influence on in vitro cellular
behavior and in vivo bone forming capacity was investigated in this study. A
simulated body fluid (SBF) test demonstrated that the incorporation of 40wt% HA
particles significantly promoted the biomineralization ability of the PU scaffolds.
Enhanced in vitro proliferation and osteogenic differentiation of the seeded MSCs
were also observed on the PU/HA composite. After being implanted subcutaneously
in a nude mice model for 8 weeks, a considerable amount of vascularized bone
tissue with initial marrow stroma development was generated in both PU and PU/
HA40 scaffold. In conclusion, the PU/HA composite is a potential scaffold for bone
regeneration applications.
3. Is it possible to fabricate a 3D scaffold by electrospinning technique and utilize
it as cell carrier to enable effective in vivo endochondral bone generation?
To further explore candidate scaffolds for endochondral bone tissue engineering
applications, in Chapter 4, a novel wet-electrospinning system, using ethanol
as collecting medium, was exploited to fabricate a cotton-like poly(lactic-coglycolic acid) (PLGA)/poly(ε-caprolactone) (PCL) scaffold. Such PLGA/PCL scaffold
consisted of an accumulation of fibers in a very loose and uncompressed manner.
Rat MSCs were seeded on these scaffolds and chondrogenically differentiated in
vitro for 4 weeks followed by subcutaneous implantation in vivo for 8 weeks. Cell
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Summary, Closing Remarks and Future Perspectives
pellets were used as a control. The glycosaminoglycan (GAG) assay and Safranin O
staining showed that the cells infiltrated throughout the scaffolds and deposited
an abundant cartilage matrix after in vitro chondrogenic priming. Histological
analysis of the in vivo samples revealed extensive new bone formation through the
remodeling of the cartilage template. In conclusion, using the wet-electrospinning
method, we are able to create a 3D scaffold in which bone tissue can be formed via
the endochondral pathway. This system can be easily processed for various assays
and histological analysis. Consequently, it is more efficient than the traditional
cell pellets as a tool to study endochondral bone formation for tissue engineering
purposes.
4. Can in vivo endochondral bone generation be achieved in various scaffold
materials?
Chapter 5 investigated the behavior of bone marrow derived mesenchymal stem
cells (MSCs) in regard to in vitro cartilage formation and in vivo bone regeneration
when combined with different three-dimensional (3D) scaffold materials, i.e.
hydroxyapatite/tricalcium phosphate (HA/TCP) composite block, polyurethane
(PU) foam, poly(lactic-co-glycolic acid)/poly(ε-caprolactone) electrospun fibers
(PLGA/PCL) and collagen I gel. To this end, rat MSCs were seeded on these scaffolds
followed by the same experimental procedures as described in Chapter 4. After
in vitro chondrogenic priming, comparable cell amounts and cartilage formation
was observed in all types of scaffolds. Nonetheless, the quality and maturity of in
vivo ectopic bone formation appeared to be scaffold/material-dependent. Eight
weeks of implantation was not sufficient to ossify the entire PLGA/PCL constructs,
albeit that a comprehensive remodeling of the cartilage had occurred. For HA/
TCP, PU and collagen I scaffolds, more mature bone formation with rich vascularity
and marrow stroma development could be observed. These data suggest that
chondrogenic priming of MSCs in the presence of different scaffold materials allows
the establishment of reliable templates for generating functional endochondral
bone tissue without using osteoinductive growth factors in vivo. The morphology
and maturity of bone formation can be dictated by the scaffold properties.
5. What is the effect of in vitro chondrogenic priming time of MSCs on in vivo
endochondral bone formation?
To enable an effective translation of this endochondral approach towards the clinic,
it is imperative to possibly shorten the in vitro procedure for the creation of a
cartilage template and meanwhile ensure sufficient bone formation. Therefore, in
Chapter 6, we aimed to investigate the influence of in vitro chondrogenic priming
time of MSCs on the in vivo endochondral bone formation, both qualitatively and
quantitatively. To this end, rat bone marrow derived mesenchymal stem cells
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were seeded on two scaffolds with distinguished features, i.e. a fibrous polymerbased PLGA/PCL scaffold or a porous ceramic-based HA/TCP composite. Next
the constructs were chondrogenically differentiated for 2, 3 and 4 weeks in vitro
followed by subcutaneous implantation for up to 8 weeks. The results showed
that, a longer chondrogenic priming time resulted in a significantly increased
amount and homogeneous deposition of the cartilage matrix on both PLGA/
PCL and HA/TCP scaffolds in vitro. In vivo, all implanted constructs gave rise to
endochondral bone formation, whereas the bone volume was not affected by a
longer priming time. An unpolarized woven bone like structure, with significant
amounts of cartilage remaining, was generated and maintained in fibrous PLGA/
PCL scaffolds, while porous HA/TCP scaffolds supported progressive lamellar-like
bone formation with mature bone marrow development.
6. What is the role of cartilage templates in mediating MSC behavior in the
process of endochondral bone formation?
Costly and lengthy process for creating cartilage templates hampers the translation
of the endochondral approach to clinic. A more comprehensive understanding of
the crosstalk between implanted cartilage template and host cells is essentially
beneficial to possibly simplify this procedure or develop substitutes for effective
bone regeneration. Therefore, in Chapter 7, the role of cartilage template in
mediating MSC behavior in the process of endochondral bone formation was
explored, as well as the possible signaling pathways involved in these biological
events. By using an indirect coculture system, our results demonstrated that
chondrogenically differentiated MSC-HA/TCP constructs (chondro-TCP) were
able to induce MSCs migration by possibly activating the IL8/CXCR2 signaling
pathway. Additionally, they could trigger a considerable elevation of signaling
factors responsible for MSCs recruitment and osteoclast stimulation, including
CCL3, CCL7, CCL19 and NF-κB. The osteogenic differentiation of MSCs was not
significantly influenced after short exposure to chondro-TCP templates.
2. Closing remarks and future perspectives
In the context of skeletal tissue development and repair, endochondral ossification
has inspired a shift of the bone tissue engineering concept from the conventional
intramembranous to the endochondral pathway. Although great challenges still
remain before this strategy can be eventually progressed into clinical application,
this thesis contributes to a further understanding of endochondral bone
regeneration. Based on the obtained results, several key conclusions can be drawn:
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Summary, Closing Remarks and Future Perspectives
(i) Compared to the use of cell pellets, chondrogenic priming of bone marrow
derived MSCs in the presence of scaffolds allows a more reliable and efficient
establishment of templates to generate functional endochondral bone tissue in
vivo, which can be achieved by using a variety of biomaterials. Among the tested
scaffold materials in this thesis, collagen I hydrogel and porous calcium phosphate
scaffold are the top choices to target a vascularized lamellar-like bone formation
with mature marrow development.
(ii) By applying a chondrogenically differentiated MSC-scaffold construct as a
cartilage template, the essential in vitro priming time can be shortened to 2 weeks
in order to generate a substantial amount of vascularized endochondral bone in
vivo. This is closely contributed by the significant formation of cartilage tissue and
expression of VEGF early to 2 weeks time point. Moreover, our data also implied
that a certain extent of in vitro chondrogenic priming of the MSCs, possibly with
even shorter time than 2 weeks, might be sufficient to ensure the progression
of endochondral differentiation of the seeded cells and enable successful bone
formation. If proved by the further investigations, this would lead to an advance
for applying the endochondral approach in the clinical practice, as a shortened
in vitro culture procedure would significantly reduce the treatment cost and
regulatory burden.
(iii) The properties of scaffold materials play an imperative role in dictating the
progression and structure of the endochondral bone formation. Among which,
biomaterial chemistry deeply affects the cell fate at early stages, but the internal
architecture of the scaffolds also plays a profound role, especially in a long term,
since it influences bone structure and its deposition kinetics. For both polymer
(e.g. PU) and ceramic (e.g. HA/TCP) based scaffolds, a confined porous structure
could concentrate collagen fibers and serve as a polarizing template within the
pores to trigger the formation of lamellar like bone. On the contrary, fibrous
scaffolds (e.g. PLGA/PCL) possessed a high level of geometrical complexity, thus
the new collagen fibers were distributed in an isotropic manner, resulting mostly
in woven bone.
(iv) After implantation, a cartilage template is able to exert a timely influence on
the recruitment of the host MSCs and the promotion of osteoclastogenesis by
activating a number of chemokine signaling related factors. Further efforts to
determine the signaling pathways involved in endochondral bone forming process
will possibly allow the development of alternative “cartilage templates” to target
bone regeneration in a simple and effective manner.
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For the ultimate success of translating the endochondral bone tissue engineering
approach to clinical practice, a number of limitations and necessary investigations
need to be addressed.
First of all, it is important to keep in mind that all discoveries on endochondral
bone tissue engineering so far emerged from cell culture and animal experiments,
which cannot be directly translated to human situation. In our studies, rat bone
marrow derived MSCs were used as the main cell source. However, for the final
clinical translation, there is no doubt that the efficacy of utilizing human cells to
generate endochondral bone has to be elucidated. Another issue that needs to
be addressed is the selection of the animal model. So far we have only tested
the in vivo bone formation in an ectopic site of immunocompromised animals.
By significantly reducing the number of variables involved in bone formation, this
model allows us to obtain an easier understanding of the tendency of cartilage
template undergoing endochondral bone formation. Still, it has to be noted
that the conclusions need to be further evidenced in a more clinically relevant
situation, i.e. in an orthotopic location of immunocompetent animals, where the
mechanical load, biochemical/inflammatory factors and endogenous progenitor/
stem cells are closely involved.
Although our studies have confirmed that the design of the scaffold can drive
the endochondral bone forming pattern, the specific parameters that dominate
this process and its mechanism could not be thoroughly concluded due to the
differences of the selected scaffolds regarding their chemical compositions and
physical structures. More complete understanding could be obtained if the scaffold
properties are more precisely controlled, e.g. all structural parameters of the
scaffolds are kept constant and meanwhile only the composition is varied, or vice
versa. In addition, larger-sized scaffolds have also to be tested in future studies,
where the cell invasion and vascularization will be more challenging. When the
interaction between the cells and scaffold materials is unraveled, we are able to
target endochondral bone regeneration in an efficient manner and on a larger
scale with the help of selecting and tailoring appropriate scaffolds.
Furthermore, because high cell density is crucial for the formation of cartilage
templates, the clinical application of the endochondral approach by using autologous
cells can be challenging due to the limits of such cell source or complexity of the
in vitro cell expansion procedure, especially when a large bone defect is involved.
At this point, the underlying mechanism about how the implanted cells induce a
specific response of the host cells and the successive tissue regeneration is still
largely unknown. Further investigations are warranted to elucidate the cellular
and molecular events involved in this process before a reliable endochondral
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Summary, Closing Remarks and Future Perspectives
therapy can be developed. More importantly, only when the crosstalk between
the implanted cells and the host cells is thoroughly elucidated, endochondral bone
tissue engineering can be further simplified and more effectively applied.
In summary, bone regeneration via the endochondral pathway requires a
deliberation of many factors, including a variety of cells, matrixes, and signals, in
an orderly temporal and spatial sequence. Although extensive efforts still have to
be devoted, endochondral bone tissue engineering holds great promise for the
final clinical application. The ultimate goal to implement this approach is not to
solely replicate the natural ossification process, but to recapitulate and extract
the essence of this process to target bone regeneration in a simple, cost-effective
and controllable way, e.g. by using scaffolds which releases various biological cues
at right timing that provoke host cell performance to target endochondral bone
formation without implantation of exogenous cells. For this to happen, a thorough
understanding of the underlying cellular and molecular mechanism of the
endochondral approach, including cell-cell and cell-matrix/material interactions,
is required, which needs collaboration among many disciplines, including cell and
molecular biology, material science, pharmacology, chemistry and medicine.
8
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Chapter 9
Samenvatting, slotopmerkingen
en toekomstperspectieven
Samenvatting, Slotopmerkingen en Toekomstperspectieven
1. Samenvatting en evaluatie van de doelstellingen
Hoewel er nog veel uitdagingen overwonnen moeten worden, is
botweefselengineering een veelbelovende aanpak om de huidige
behandelmethoden via autologe of allogene bottransplantatie te vervangen. De
standaardprocedure die gebruikt wordt bij botweefselengineering is het laden van
osteogene voorlopercellen op biomaterialen die de differentiatie van cellen in een
specifieke richting sturen, gevolgd door de implantatie van deze constructen in
het lichaam. Een dergelijke aanpak beoogt niet alleen het beschadigde weefsel
te vervangen, maar induceert en dirigeert ook het genezingsproces om tot een
volledige genezing van het aangedane weefsel te komen. De huidige methoden
om bot te genereren zijn grotendeels gericht op intramembraneuze ossificatie
(d.w.z. botvorming via osteoblast differentiatie), vooralsnog is de werking van deze
methode echter nog niet op overtuigende wijze resultaat getoond bij mensen.
Recentelijk heeft een nieuwe aanpak, gebaseerd op endochondrale ossificatie, de
aandacht getrokken. Deze aanpak richt zich op het vormen van bot door middel van
kraakbeen. Vergeleken met osteogeen gedifferentieerde cellen zijn chondrocyten
goed aangepast aan hypoxische condities met daarbij een beperkt aanbod
van nutriënten. Bovendien produceren chondrocyten tijdens de differentiatie
aanzienlijke hoeveelheden angiogene factoren. Door deze eigenschappen heeft
de endochondrale aanpak een groot voordeel bij het levensvatbaar houden van
de geïmplanteerde cellen en het genereren van gevasculariseerd botweefsel.
Het onderzoek en de beschikbare kennis over de endochondrale methode staan
echter nog in de kinderschoenen.
Het onderzoek beschreven in dit proefschrift heeft als doel om botvorming via
de endochondrale methode te optimaliseren en zodoende een hoge kwaliteit
botherstel te bereiken. Meer specifiek werd daarbij in verschillende studies de
toepasbaarheid van verschillende scaffold materialen, optimale in vitro kweektijd
en het mechanisme van de omzetting van kraakbeen naar bot onderzocht.
Hoofdstuk 1 geeft een algemene inleiding over botweefselengineering, daarnaast
bevat dit hoofdstuk de doelstellingen van dit proefschrift. Elk volgend hoofdstuk
beschrijft een deelonderzoek. In deze samenvatting worden de vraagstellingen
zoals beschreven in het eerste hoofdstuk in opeenvolgende volgorde behandeld.
1. Wat is de huidige stand van de techniek van het combineren van stamcellen
en biologisch afbreekbare materialen voor botherstel via de endochondrale
route?
Dankzij verbeterde kennis over de bronnen van stamcellen, samen met een brede
vooruitgang in de ontwikkeling van materialen, lijkt het erop dat “engineering”
in toenemende mate toepasbaar is voor het regenereren van grote botdefecten.
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Hoofdstuk 2 gaat eerst dieper in op de eigenschappen en mechanismen van
botweefsel en osteogenese, om zo een fundamenteel begrip van de betrokken
processen te verkrijgen. Vervolgens worden de in de literatuur beschreven
onderzoeken over de toepassing van biomaterialen en celcultuur strategieën om
endochondrale botvorming te verkrijgen toegelicht. Van verschillende type cellen,
waaronder chondrocyten, MSCs en ESCs, is het bekend dat ze het vermogen bezitten
om endochondrale differentiatie te ondergaan, met of zonder specifieke stimulatie.
Verschillende biomaterialen, met een structuur en samenstelling die nauwkeurig
de aard kraakbeen of botweefsel nabootsen, zijn potentiële kandidaten voor de
vorming van osteochondrale weefsels. Natuurlijke of synthetische polymeren en
keramische materialen zijn daarbij de belangrijkste kandidaten. De mogelijkheid
tot het genereren van gevasculariseerd bot is aangetoond via de endochondrale
methode, hoewel er standaardisatie en optimalisatie nodig zal zijn om voor een
efficiënte omzetting van kraakbeen naar botweefsel te zorgen. Voordat een
klinische test uitgevoerd kan worden moeten bovendien nog een aantal zaken
uitgezocht worden, waaronder het juiste tijdstip voor de implantatie, de functie
en het lot van de geïmplanteerde cellen.
2. Is het poreuze alifatische polyurethaan/hydroxyapatiet composiet een
geschikte scaffold voor botweefselengineering toepassingen?
Biomaterialen met als doel te functioneren als ondersteunende structuur voor
osteogene cellen spelen een centrale rol in botweefselengineering. In hoofdstuk
3 beschrijven we hoe een alifatische polyurethaan (PU) schuim gesynthetiseerd
werd met behulp van niet-giftige bestanddelen. Via deze procedure werd een
uniforme poreuze structuur gecreëerd, waarbij de holtes in de structuur onderling
verbonden waren. Bovendien werden hydroxyapatiet (HA) deeltjes aan dit
proces geïntroduceerd om de bioactiviteit van de PU matrix te verhogen. Om de
biologische prestaties van deze PU scaffolds te evalueren, werd hun invloed op het
cellulaire gedrag in vitro en de bot vormende capaciteit in vivo onderzocht. Een
test met “simulated body fluid” (SBF) toonde aan dat het inbedden van 40 gew%
HA deeltjes het vermogen tot biomineralisatie significant verhoogde. Verbeterde
cel proliferatie en osteogene differentiatie van de gekweekte MSCs werden
waargenomen op de PU/HA composiet. Na subcutane implantatie in een naakte
muis model voor 8 weken, werd een aanzienlijke hoeveelheid gevasculariseerd
botweefsel en beenmerg ontwikkeling gegenereerd bij zowel de PU als de PU/
HA40 scaffolds. De conclusie is dat het PU/HA composiet materiaal veel potentiee
heeft voor botregeneratieve toepassingen.
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Samenvatting, Slotopmerkingen en Toekomstperspectieven
3. Is het mogelijk om een 3D-scaffold te fabriceren via elektrospinning en deze te
gebruiken als een drager materiaal voor de cellen om in vivo endochondrale
bot generatie effectief mogelijk te maken?
Om verder te onderzoeken welke scaffolds geschikt zijn voor endochondrale
botweefselengineering toepassingen, is in hoofdstuk 4 beschreven hoe een
nieuwe natte-elektrospin techniek werd ingezet. Hierbij werd gebruik gemaakt
van ethanol als medium om de vezels in te verzamelen. Via deze methode werd
een katoenachtige poly (melkzuur-co-glycolzuur) (PLGA)/poly (ε-caprolacton)
(PCL) scaffold gefabriceerd, bestaande uit een opeenstapeling van vezels op
een zeer losse ongecomprimeerde wijze. Op deze scaffolds werden vervolgens
ratten beenmergcellen gezaaid en deze werden chondrogeen gedifferentieerd
in vitro gedurende 4 weken, gevolgd door subcutane implantatie in vivo voor
8 weken. Daarbij werden cel pellets gebruikt als controle. De analyse van
glycosaminoglycaan (GAG) en safranine O kleuring toonde aan dat na in vitro
chondrogene priming de cellen de gehele matrix hadden geïnfiltreerd en daarbij
een overvloedige kraakbeenmatrix hadden gedeponeerd. Histologische analyse
van de in vivo constructen demonstreerde uitgebreide nieuwe botvorming door
hermodellering van het kraakbeen. Samenvattend, gebruik makend van de natte
electrospinning methode, kunnen wij een 3D scaffold creëren die helpt om via
de endochondrale route bot te vormen. Dit systeem leent zich uitstekend voor
analyse met verschillende assays en histologische analyse. Daarom is dit systeem
efficiënter voor het bestuderen van weefselregeneratie via de endochondrale
methode dan het gebruik van de traditionele cel paletten.
4. Kan via de endochondrale route bot worden gevormd in verschillende
materialen?
In hoofdstuk 5 wordt het gedrag van mesenchymale stamcellen (MSCs) uit het
beenmerg onderzocht met betrekking tot de in vitro vorming van kraagbeen en
in vivo botregeneratie in combinatie met verschillende driedimensionale (3D)
materialen, waaronder: hydroxyapatiet/tricalciumfosfaat (HA/TCP), polyurethaan
(PU) schuim, poly (melkzuur-co-glycolzuur)/poly (ε-caprolacton) vezels (PLGA/
PCL, bereid via electrospinning) en collageen-I gel. Hiertoe werden ratten MSCs
gezaaid op deze scaffolds gevolgd door dezelfde experimentele procedure als
beschreven in hoofdstuk 4. Vergelijkbare hoeveelheden cellen en kraakbeen
werden waargenomen na in vitro chondrogene priming voor alle soorten scaffolds.
Toch leek de kwaliteit en de maturatie van het in vivo ectopische gevormde bot
materiaal afhankelijk. Een implantatietijd van acht weken was niet voldoende om
de volledige PLGA/PCL constructen te ossificeren, maar het kraakbeen was wel
duidelijk geremodelleerd. Voor de HA/TCP, PU en collageen I scaffolds, werd een
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meer volwassen botvorming met rijke vascularisatie en beenmerg ontwikkeling
waargenomen. Deze gegevens suggereren dat chondrogene priming van MSCs in de
aanwezigheid van verschillende scaffold materialen het genereren van functioneel
botweefsel via de endochondrale route zonder osteoinductie groeifactoren in vivo
mogelijk maakt. De morfologie en de maturatie van botvorming worden mede
gedicteerd door de eigenschappen van het scaffold materiaal.
5. Wat is het effect van in vitro chondrogene priming van de MSCs op in vivo
endochondrale botvorming?
Om een effectieve behandeling van deze endochondrale aanpak klinisch mogelijk
te maken, is het noodzakelijk om de in vitro procedure voor het fabriceren van een
kraakbeen substraat zo kort mogelijk te houden en ondertussen te zorgen voor
voldoende botvorming. In hoofdstuk 6 is daarom onderzocht wat de invloed is
van de in vitro chondrogene primings tijd van MSCs op de in vivo endochondrale
botvorming, zowel kwalitatief als kwantitatief. Hiervoor werden van ratten MSCs
gezaaid op twee scaffolds met verschillende kenmerken; een vezelachtige PLGA/
PCL polymeer of een poreuze keramische, HA/TCP composiet. Vervolgens werden
de constructen voor 2, 3 of 4 weken chondrogeen gedifferentieerd in vitro, gevolgd
door subcutane implantatie tot 8 weken. De resultaten toonden aan dat een langere
chondrogene primings tijd resulteerde in een significant verhoogde hoeveelheid
afzetting van kraakbeenmatrix voor zowel de PLGA/PCL en HA/TCP materialen
in vitro. In vivo, gaven alle geïmplanteerde substraten uiteindelijk aanleiding tot
endochondrale botvorming, terwijl het bot volume niet werd beïnvloed door een
langere priming tijd. Een ongepolariseerde gewoven bot structuur, waarbij nog
aanzienlijke hoeveelheden kraakbeen zichtbaar waren, werd gegenereerd bij de
vezelachtige PLGA/PCL scaffolds, terwijl de poreuze HA/TCP scaffolds lamellairachtige botvorming met een rijpere beenmerg ontwikkeling ondersteunden.
6. Op welke wijze beïnvloeden kraakbeen substraten het gedrag van MSCs
tijdens het endochondrale proces?
Een kostbaar en langdurig proces voor het maken van kraakbeen substraten
belemmert de klinische haalbaarheid van de endochondrale methode. Een
completer begrip van de communicatie tussen het geïmplanteerde kraakbeen
substraat en de cellen van de ontvanger is noodzakelijk om deze methode verder
te verbeteren. Daarom is in hoofdstuk 7 de invloed van het kraakbeen substraat
op het gedrag van MSCs tijdens endochondrale botvorming onderzocht en werd
daarnaast onderzocht welke biologische signalen mogelijk betrokken zijn bij dit
proces. Door gebruikt te maken van een indirect cocultuur systeem, toonden
onze resultaten aan dat chondrogeen gedifferentieerde MSC-HA/TCP substraten
(chondro-TCP) migratie van MSCs mogelijk veroorzaken door het activeren van het
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Samenvatting, Slotopmerkingen en Toekomstperspectieven
IL8/CXCR2 signaal transductie route. Daarnaast werd een aanzienlijke verhoging van
de signalerings factoren, zoals CCL3, CCL7, CCL19 en NF-kB, die verantwoordelijk
zijn voor het rekruteren van MSCs en het stimuleren van osteoclasten gevonden.
De osteogene differentiatie van MSCs werd niet significant beïnvloed na korte
blootstelling aan de chondro-TCP templates.
2. Slotwoord en toekomstperspectieven
In het kader van de ontwikkeling van het skelet en weefselherstel, inspireert de
endochondrale ossificatie een verschuiving in de botweefselengineering van de
conventionele intramembraneuze naar de endochondrale route. Hoewel er nog
steeds grote uitdagingen voordat deze strategie uiteindelijk kan worden toegepast
in de kliniek, draagt dit proefschrift bij aan een beter begrip van het endochondraal
botherstel. Op basis van de verkregen resultaten kan een aantal belangrijke
conclusies worden getrokken:
(i) In vergelijking met het gebruik van celpellets, biedt chondrogene priming
van MSCs uit het beenmerg in de aanwezigheid van materialen veel voordelen
waaronder het creëren van een betrouwbaar en efficiëntesubstraat dat via de
endochondraal route botweefsel kan genereren. Dit kan worden bereikt worden
met een verscheidenheid aan biomaterialen. Onder de in dit proefschrift geteste
materialen geven collageen-I hydrogelen en poreus calciumfosfaat de beste
resultaten om een gevasculariseerd lamellair-achtige botvorming met volwassen
beenmerg ontwikkeling te verkrijgen.
(ii) Door gebruik te maken van een construct met chondrogeen gedifferentieerde
MSCs als een kraakbeen substraat, kan de essentiële in vitro priming tijd om
een aanzienlijke hoeveelheid gevasculariseerde bot te genereren verkort
worden tot 2 weken. Dit was duidelijk zichtbaar door de aanzienlijke vorming
van kraakbeenweefsel en de expressie van VEGF binnen 2 weken. Bovendien,
suggereren de gegevens ook dat eventueel een nog kortere primings tijd dan 2
weken, kan volstaan voor voldoende chondrogene differentiatie van de gezaaide
cellen en uiteindelijk dus botvorming. Als dit inderdaad bewezen wordt door de
verdere onderzoeken, zou dit een groot voordeel zijn voor de toepassing van
de endochondrale aanpak in de klinische praktijk, omdat de verkorte in vitro
cultuur tijd de kosten en administratieve lasten van de behandeling aanzienlijk
verminderen.
(iii) De eigenschappen van scaffold materialen spelen een belangrijke rol in het
dicteren van de progressie en de structuur van de endochondrale botvorming.
De chemische eigenschappen van het biomateriaal spelen vooral een rol op de
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korte termijn, terwijl de interne structuur ervan meer invloed heeft op de lange
termijn aangezien het de botstructuur en de depositie kinetiek beïnvloedt. Voor
zowel polymeren (bijv. PU) als keramiek (bijv. HA/TCP) gebaseerde scaffolds, kan
een structuur met kleine poriën collageenvezels concentreren en dienen als een
polariserend sjabloon in de poriën en zo de vorming van lamellair bot induceren.
Vezelige scaffolds (bijv. PLGA/PCL) daarentegen bezitten een hoge geometrische
complexiteit, waardoor de nieuwe collageen vezels zich op een isotrope manier
verspreiden, wat meestal leidt tot gewoven bot.
(iv) Na implantatie is een kraakbeen substraat in staat om invloed uit te oefenen
op de MSCs van de ontvanger en zo osteoclast vorming te induceren door het
activeren van een aantal aan chemokines gerelateerde factoren. Verdere studies
om meer begrip te krijgen voor de signaal routes betrokken bij endochondrale
botvorming, kunnen mogelijk de ontwikkeling van alternatieve “kraakbeen
substraten” naar botregeneratie vereenvoudigen en efficiënter maken.
Om de endochondrale aanpak voor botweefselengineering succesvol toe te passen
in de klinische praktijk, moeten nog een aantal hordes genomen worden. Allereerst
is het van belang om in het achterhoofd te houden dat alle resultaten m.b.t. de
endochondrale methode zijn voortgekomen uit celkweek en dierproeven. Deze
resultaten zijn dus niet direct te vertalen naar menselijke situaties. In onze studies
bijvoorbeeld, werden MSCs gebruikt verkregen uit het beenmerg van ratten. Er is
geen twijfel over dat voor de uiteindelijk klinische toepassing de effectiviteit van
het gebruik van menselijke cellen opnieuw zal moeten worden opgehelderd. Een
ander probleem dat moet worden aangepakt is de selectie van het diermodel.
Tot dusver hebben we alleen de in vivo botvorming op een ectopische plaats
van immunogecompromitteerde dieren getest. Dit model werkt door het aantal
variabelen betrokken bij botvorming aanzienlijk te verminderen, waardoor er
gemakkelijker er een simpelere vertaling is van de overgang van kraakbeen
substraat naar bot . Daarom moeten de uitkomsten getoetst moeten worden in een
klinisch relevante situatie, dwz een orthotopische locatie in immunocompetente
dieren, waarbij de mechanische belasting, biochemische/ontstekingsfactoren en
stamcellen nauw betrokken zijn.
Hoewel onze studies hebben bevestigd dat de structuur van de scaffolds het
patroon van botvorming kunnen beïnvloeden, is het moeilijk een uitspraak te doen
over het precieze mechanisme daarvan. De chemische en fysische eigenschappen
van de gebruikte scaffolds verschilden daarvoor teveel. Om dit beter te kunnen
onderzoeken is het nodig om een serie materialen te ontwikkelen waarbij telkens
slechts een parameter (fysisch, chemisch, enz) gevarieerd wordt. Daarnaast
is het belangrijk om grotere scaffolds te testen, omdat de invasie van cellen en
166
Samenvatting, Slotopmerkingen en Toekomstperspectieven
vascularisatie een grotere uitdaging zal zijn. Wanneer de interactie tussen de cellen
en scaffold materialen is ontrafeld, is het mogelijk om op efficiëntere en grotere
schaal endochondrale botregeneratie toe te passen mede door het gebruikt van
afgestemde materialen.
Een hoge celdichtheid is cruciaal is voor de vorming van de kraakbeen substraten,
dit kan de klinische toepassing van de endochondrale benadering met autologe
cellen uitdagend maken. De beschikbaarheid van deze cellen kan dan een
beperkende factor worden, zeker wanneer het om een groot botdefect gaat. Op dit
moment is het onderliggende mechanisme van het door de geïmplanteerde cellen
induceerde reactie op de gastheercellen en de opeenvolgende weefselregeneratie
nog grotendeels onbekend. Verder onderzoek is nodig om de cellulaire en
moleculaire gebeurtenissen die betrokken zijn bij dit proces op te helderen, voordat
een betrouwbare endochondrale therapie kan worden ontwikkeld. Belangrijker
nog, is het begrijpen van de communicatie tussen de geïmplanteerde cellen en
de gastheercellen, alleen wanneer dit tot in detail begrepen is kan endochondrale
botweefselengineering verder worden vereenvoudigd en efficiënter worden
toegepast.
Samengevat, botherstel via de endochondrale route vereist begrip van vele
factoren, waaronder een groot aantal verschillende cellen, matrices en signalen.
Hoewel er nog grote inspanningen verricht moeten worden, houdt endochondrale
botweefselengineering een grote belofte in voor klinische toepassing. Het
uiteindelijke doel van deze aanpak is niet alleen het kopiëren van het natuurlijke
ossificatieproces, maar ook het recapituleren van de essentie van dit proces op
een simpele en kosten efficiënte manier. Dit kan bereikt worden door scaffold
materialen die op het juiste moment de juiste biologische signalen afgeeft om
endochondrale botvorming te verkrijgen zonder implantatie van cellen. Om dit
te kunnen realiseren is een grondig begrip nodig van de onderliggende cellulaire
en moleculaire mechanismen van de endochondrale route, met inbegrip van
cel-cel en cel-matrix/materiaal interacties. Dit vereist de samenwerking van
vele disciplines, met inbegrip van cel- en moleculaire biologie, materiaalkunde,
farmacologie, chemie en geneeskunde.
9
167
Acknowledgements
List of publications
Curriculum Vitae
Acknowledgements
Life is like a box of chocolates. You never know what you are going to get. Before
2010, I could not imagine such a long journey in a country so far away from my
family and friends. But now, to me, Netherlands has become not only the witness
for my growth during the PhD, but also a sweet home where I discovered and
harvested so much beauty and love. While writing the last chapter of my thesis,
I cannot help bringing back a lot of nice and meaningful memories I gathered in
the past four years. The help, guidance, trust, support and love from the people
I have stayed with, certainly made this journey so unique, colorful, and exciting.
Although it is not possible to mention everyone by name, I would like to specifically
acknowledge the following people.
First and foremost, I would like to express my sincere gratitude to my supervisor,
Prof. dr. John Jansen. Dear John, thank you for giving me the opportunity to
experience such a wonderful journey in your department. You opened the door of
tissue engineering realm for me and I have learned so much from you in the past
years, not only from your broad knowledge and solid expertise in biomaterial field,
but also from your way of critical thinking and highly efficient working attitude.
Besides the research, the experience of working with you in the clinic has also
being very enjoyable and enlightening for me. Your enthusiasm for work, skillful
and elegant clinical operations, as well as the great considerations and respect to
the patients has set a superb model for me. All of these will absolutely encourage
me and constitute a valuable asset in my future career.
I am also deeply grateful to my co-supervisor, Dr. Fang Yang. Dear Fang, I am so lucky
to have you around in the past few years. To me, you are not only a supervisor who
gave solid supports and trust for me and my experiments, but also a friend, a sister,
who are always willing to help me, and share the joyful or difficult moments in my
life. Your expertise in biomaterial and critical attitude has helped me make a lot
of progress in research and writing. Your warm-hearted character and optimistic
attitude towards work and life always encouraged and inspired me. I would never
adapt to the Dutch life so quickly and pleasantly without your guidance on how to
understand and appreciate the difference between western and eastern culture.
In the past 4 years, I could not count how many joyful moments we have shared
together: Chinese dinners, table games, BBQ, Karaoke, sports, travelling,... I
sincerely thank you, and also Ronald, for the warm company during my PhD. This
friendship will never fade in the future.
I would like to further express my gratitude to my other co-supervisor, Dr. Sanne
Both. Dear Sanne, you lead me to the field of cell culture from the very beginning,
and it turned out that I enjoyed a lot working with those tiny and energetic creatures
170
Acknowledgements
during my PhD. Your solid knowledge in cell culture and broad experiences in many
biological techniques provided helpful guidance for my experiments. I appreciate
your great patience in teaching; your vivid explanations always made me easily
get the points. You showed a good example for me, not only being precise and
collaborative on work, but also being confident, open-minded and adventurous in
life. It has been very pleasant to be your colleague and friend. Thank you for all of
your supports, encouragements and trust during my PhD; and also thank you for
taking me to all those wonderful trips: paragliding, Christmas night in Gouda, and
so on.
I would also like to extend my appreciation to Prof. Yining Wang, Prof. Gerjo van
Osch and Prof. Pieter Buma. Dear Prof. Wang, I was lucky to be your student during
my master study, and thank you for the continuous encouragements in the last
few years. The knowledge and working attitude I learned from you have firmly
supported me pursuing the goal of PhD. Dear Prof. van Osch, I would never finish
my studies so smoothly without the help from you and also the technical support
from your lab in Rotterdam. Thank you for all those valuable and efficient advices
for my studies and articles. Dear Prof. Buma, it was my pleasure to have you as my
mentor. Thank you for the inspiring discussions about my studies.
My appreciation also goes to my colleagues Dr. Joop Wolke, Dr. Frank Walboomers,
Dr. Jeroen van den Beucken, Dr. Sander Leeuwenburgh, and former colleague
Dr. Marco Lopez-Heredia. Your sharp insight and solid expertise in biomaterials
field strongly supported my research. Thank you for your advice in all kinds of
discussions and group meetings. Dear Joop, I admire your critical attitude to
research and the warm heart whenever I or anyone else needs your help or advice.
Dear Frank, it has been always a pleasure to talk with you. I was inspired not only
by your humorous stories, but also your positive attitude towards life and great
passion for the family. Dear Jeroen, I was always impressed by your high working
efficiency. Thank you for all the laughers you brought to us and also for the nice
company during the trip in Greece. Dear Sander, I learned from you to be generous
and modest as a person, precise and efficient in work, passionate and adventurous
for life. Thank you for every encouragement after my presentations. Dear Marco,
I would never forget the pleasant time working with you during my first trip in
Nijmegen as an exchange student. You were such a patient teacher leading a
novice like me approaching science, and also a good friend sharing all kinds of
stories in life. Your earnestness and diligence towards work has certainly set an
excellent example for me and encouraged me to start my PhD journey.
My work would not be completed so smoothly without the help from all “superhero” technicians. I would like to express my gratitude to Natasja, Vincent,
171
Monique, Martijn and Rene for your patience and being always there helping us.
Special thank goes to Natasja, not only for your great support in my work, but
also the pleasant trips we went together. I will always remember the beautiful
stainings we worked out together, and the joyful swimming lessons. I would also
like to thank Wendy Koevoet and Nicole Kops from Erasmus MC, for the technical
supports and advices for my experiments.
My acknowledgement also goes to all the people from CDL who helped me arrange
and perform animal studies. In particular, I would like to thank Debby, Jeroen and
Daphne for your friendly assistance.
Many people in dental school have given me all kinds of help during my PhD.
Dear Kim, Henriëtte, Monique and Vera, I appreciate your professional and
patient support in the past years. Dear Rachel, thank you for the friendly guide
in the clinic. I also want to acknowledge Dr. Ruud Kasten, Dr. Aukje Bouwman,
and Dr. Ewald Bronkhorst. Dear Ruud, I would never have the chance to start my
journey in Nijmegen without your invitation in 2009. Thank you for all your warm
encouragements in the past years. Dear Aukje and Ewald, thank you for your
valuable advices and help for my experiments.
My stay in Netherlands would have never been so pleasant and enjoyable without
my wonderful colleagues in department of Biomaterials. Thank you all for creating
such a warm and friendly environment. Thank you all for sharing with me the
knowledge in experiments and every story in life. Because of you all, working was
never just working; I enjoyed every cheerful and joyful day with my friends here.
My first heartfelt appreciation goes to Xiangzhen and Na. Dear Xiangzhen, I can
still vividly recall the very first day we flew thousands kilometers away from home
and landed together in this new place. In the past 4 years, we took care of each
other as colleagues, best friends and sisters. You showed me great character of
persistence, honesty, courage and strong mind. I cannot count how many times I
turned to you whenever I met happy, depressed, confusing or just bored moments
in my life. Thank you for always being there listening and supporting me. Although
we may not see each other often after the PhD, I will always cherish the fantastic
trips we have made, the beautiful handcraft shops we have stopped, and the
childish dreams we have talked.
Dear Na, from bachelor to PhD, our friendship grows tighter and tighter. You are
the witness for every step I made forward, and also such a loyal friend that always
supports me with your best efforts. Thank you and also Zengping for hosting so
many nice activities and delicious dinners. Thank you for every encouragement
and wish you made for my work and life. Your homelike company made me never
feel lonely even far away from my own family.
172
Acknowledgements
My PhD journey could never be so cheerful without my lovely (ex)officemates
from all different countries: Eva, Alexey, Kemal, Daniel, Kambiz, Mani, Rosa, Rania
and Floor. Dear Eva, it has been such a pleasant journey for me to have you around
from the beginning of my PhD. Your passion and sincerity to life, friends and family,
always inspired and moved me. I will never forget our fantastic adventures in US,
and so many other moments we have shared together; no matter happy or sad,
we grow stronger and more confident together towards our life and future.
Dear Alexey, you are such an easy-going and helpful person. I was lucky to have
you around. Your optimism and generous help always calmed me down when I
met difficulties. Your humors and stories made me never feel bored staying in the
office. I will always remember the countless laughers we have had, the bets we
have made, and also the cookies you have stolen from me. :) Thanks also go to
Thordis for volleyball, BBQ and many other nice activities together.
Dear Kemal and Daniel, from 2nd to 1st floor, we were always united in the same
office. I like the office time that we were sharing those serious or non-serious
knowledge; and I also keep a lot of happy memories from our hanging-out time
after work. Thank you, Kemal, for all kinds of warm invitations and for letting
me “beat” you in pingpang. Dear Kambiz and Mani, you created such a relaxing
and “niiicee” atmosphere for me during the last period of my PhD. Your humors
always released me from the stress of writing. Thank you for the friendly advices
whenever I turned to you. Dear Rosa, Rania and Floor, thank you for all the warm
encouragements and nice chats. I always miss our old cozy office and all the joyful
moments filled there.
I never feel alone also because I have a big Chinese family in the department
supporting me at different periods of my PhD. Dear Huanan, thank you for your
company playing all kinds of sports with me, searching dinners with me in the
canteen after work, and having those relaxing chats. You set a good example for
Chinese guys in our department: diligent in work, optimistic in life, passionate for
family, but maybe not your cooking skills. :) Dear Jinling, you were my greatest
helper from the beginning of my PhD and I learned a lot from your optimism and
persistence. I can still vividly recall the days we explored the experiments together
and encouraged each other, which made me feel more confident and supported.
I also cherish all the good time we shared in the office, in the sports center and
in many other corners of our life. Dear Wei and Jie, you were always so helpful
whenever I discussed the experiment or the daily life with you. I also miss those
girls’ talks and many joyful activities we spent together. Dear Jiankang, Xinjie, Yang
and Yue, staying with you guys makes me always feel younger and more energetic.
I will miss a lot the delicious meals you have cooked.
173
It would not be a complete memory for the days in Biomaterials, if I missed the
names of many friendly colleagues that have been or still are in this big family:
Dear Manuela and Ruggero, you started my exploration of this city; it was much
fun to hang out with you. Thank you for all those warm encouragement and advice
for my work and life. Dear Paula and Astghik, it has been always enjoyable to
talk with you, and I learned a lot from your cooking lessons. Thanks also to dear
colleagues Ljupcho, Matilde, Edwin, Hamdan, Cristina, Jan Willem, Bart, Daniël,
Claire, Antonio, Pedro, Simone, Reza, Ana, Michele, Take, Paulo, Jing, Weihua,
Hongbin, and “neighbor” colleagues Aysel and Niels. Thank you for the company
along the journey.
Besides work, I also shared a lot of cheerful moments with my housemates. Dear
Ineke, I learned a lot from you to discover the simple beauty from life. Thank you
for taking care of me when I was sick, and the nice talks during our dinners. Dear
Lie and Yi, it was such a warm feeling when we built a home together. I will always
remember the days we spend together in Vossenlaan.
The company and support from friends outside the department also made my PhD
journey more colorful. Thank you, Weiwei, Xuan, Shuo, Junru, Gu Yan, Sun Yan,
Yaping, Shanshan, Xuehui, Xinhui, Jiabo, Siji, Haiteng, Shuiquan, Junxiao, Jinge,
Daixiao, Bronte, Tina, Andrew, Steven, Johnny, Esther, Richard, Bobby, Steffie, and
many others. Although the time we spent together might not be so long, but the
nice memory will always stay. I could also not forget the warm heart and support
from my dear friends in China: Sun Wei, Xiling, Shiqian, Zhejun, Huiting, Yuanyuan,
and Manli. Although we are separated from such a long distance and heading for
different life directions, our friendship will never change.
I am grateful to the Nijhuis members. Dear Marianne, Henk, Carolien, Jelle and
Marius, thank you for supporting me and taking me as your family. This always
warms me.
To my beloved parents and grandparents, I miss all of you! Your unconditional and
endless love, trust and support is the most precious fortune I could have in my life.
亲爱的老爸老妈，感谢你们三十年来不辞辛苦的养育我，教育我。你们无
私的爱和信任为我撑起了一片温暖的天空，让我勇敢的飞的更高更远去寻
找自己的梦想；你们无微的呵护和牵念是一棵繁茂的大树，时时刻刻守护
着我为我遮风避雨。女儿一直漂泊在外，很愧疚未能陪在你们身边，唯有
在心里默默的感激你们对我的宽容和体谅。也许永远也比不上你们为我付
出的三十年，但是女儿希望能用后面的三十年，五十年，无论多长，报答
你们，也成为你们的树，让你们依靠。还有我亲爱的外爹外奶，忘不了你
们为我忙碌的身影，还有你们对我一次次的叮咛和挂念，祝愿你们一直健
康平安。
174
Acknowledgements
I owe my last words to you, Arnold, my dearest colleague, friend, and love. To me,
the biggest achievement of the PhD is not about the completion of this thesis;
having met you and building a life together with you was the most meaningful and
beautiful gift I got. Thank you for catching me into your world, for spoiling me with
your powerful love, for supporting me unlimitedly in my work and life. I enjoyed
every single second I spent with you along our journey, and I know this journey
will never end.
175
List of publications
Related to this thesis:
1. Yang, W.*, Bongio, M.*, Yang, F., Both, S.K., Leeuwenburgh, S.C.G., van den
Beucken, J.J.J.P., and Jansen, J.A., Combining Osteochondral Stem Cells and
Biodegradable Hydrogels for Bone Regeneration. (Stem Cells and Bone Tissue,
2013. CRC Press: p. 326-348.)
2. Yang, W., Yang, F., Wang, Y., Both, S.K., and Jansen, J.A., In vivo bone generation
via the endochondral pathway on three-dimensional electrospun fibers. (Acta
Biomaterialia, 2013. 9(1): p. 4505-12.)
3. Yang, W., Both, S.K., van Osch, G.J.V.M., Wang, Y., Jansen, J.A., and Yang, F.,
Performance of different three-dimensional scaffolds for in vivo endochondral
bone generation. (European Cells & Materials, 2014. 27: p. 350-364.)
4. Yang, W., Both, S.K., Zuo, Y., Tahmasebi Birgani, Z., Habibovic, P., Li, Y., Jansen,
J.A., and Yang,F., Biological evaluation of porous aliphatic polyurethane/
hydroxyapatite composite scaffolds for bone tissue engineering. (Journal of
Biomedical Materials Research Part A, submitted.)
5. Yang, W., Both, S.K., van Osch, G.J.V.M., Wang, Y., Jansen, J.A., and Yang, F., Effects
of in vitro chondrogenic priming time of bone marrow derived mesenchymal
stem cells on in vivo endochondral bone formation. (Acta Biomaterialia,
submitted)
6. Yang, W., Both, S.K., Jansen, J.A., and Yang, F., Roles of cartilage templates
in mediating mesenchymal stem cell behavior in endochondral bone tissue
engineering (manuscript in preparation).
Other publications:
1. Yang, W.*, Yan, X.Z.*, Yang, F., Kersten-Niessen, M., Jansen, J.A., and Both,
S.K., Effects of Continuous Passaging on Mineralization of MC3T3-E1 Cells with
Improved Osteogenic Culture Protocol. (Tissue Engineering Part C Methods,
2014. 20(3): p. 198-204.)
2. Sariibrahimoglu, K.*, Yang, W.*, Leeuwenburgh, S.C.G., Yang, F., Wolke, J.G.,
Zuo, Y., Li,Y., Jansen, J.A., Development of porous polyurethane/strontiumsubstituted hydroxyapatite composites for bone regeneration. (Journal of
Biomedical Materials Research Part A, 2014. accepted)
3. Lopez-Heredia, M.A., Yang, W.*, Sariibrahimoglu, K.*, Bohner, M., Yamashita,
D., Kunstar, A., van Apeldoorn, A.A., Bronkhorst, E.M., Felix Lanao, R.P.,
176
List of publications
Leeuwenburgh, S.C.G., Itatani, K., Yang, F., Salmon, P., Wolke, J.G., and Jansen,
J.A., Influence of the pore generator on the evolution of the mechanical
properties and the porosity and interconnectivity of a calcium phosphate
cement. (Acta Biomaterialia, 2012. 8(1): p. 404-14.)
4. Both, S.K., Yang, W., Yang, F., and Jansen, J.A., The influence of 3D scaffolds in
endochondral bone formation. (Journal of Tissue Engineering and Regenerative
Medicine, 2012. 6: p. 33.)
5. Yeatts, A.B., Both, S.K., Yang, W., Alghamdi, H.S., Yang, F., Fisher, J.P., and Jansen,
J.A., In vivo bone regeneration using tubular perfusion system bioreactor
cultured nanofibrous scaffolds. (Tissue Engineering Part A, 2014. 20(1-2): p.
139-46.)
6. Cai, X., Yang, F., Yan, X.Z., Yang, W., Yu, N., Oortgiesen, D.A.W., Wang, Y., Jansen,
J.A., and Walboomers, X.F., Influence of bone marrow-derived mesenchymal
stem cells differentiation in periodontal regeneration: an in vivo study. (Journal
of Clinical Periodontology, submitted)
* These authors contributed equally.
177
Curriculum Vitae
Wanxun Yang (杨婉洵) was born on June 15, 1985 in
Bengbu (China). In September 2003, she was enrolled
in Wuhan University, majoring in Dentistry with sevenyear schooling system. She received basic medicine
education, resident training as a general dentist,
and postgraduate training as a prosthodontist in the
Department of Prosthodontics under the supervision
of Prof. Yining Wang. In 2009, she performed a twomonth internship in the Department of Biomaterials
at Radboud University as an exchange student. In
June 2010, she was conferred the Master Degree of
Dental Medicine with Outstanding Graduate Award
of Wuhan University. The same year, she obtained the national medical license as
a dental practitioner.
In October 2010, she was supported by Chinese Scholarship Council (CSC) and
started her PhD study in the Department of Biomaterials at Radboud University
Medical Center, under the supervision of Prof. John A. Jansen. Her project
mainly focused on combining biomaterials and pluripotent stem cells for bone
regeneration through the endochondral ossification pathway, among which
two studies were in collaboration with Department of Orthopedics, Erasmus
Medical Center Rotterdam. The results of her PhD studies are described in this
thesis and published as separated research articles in scientific journals. Part of
the results were also presented in several international conferences, including
International Bone-Tissue-Engineering Congress (Hannover, 2011), International
Society of Differentiation Conference (Amsterdam, 2012), Tissue Engineering and
Regenerative Medicine International Society (TERMIS) EU Congress (Istanbul,
2013), and the 5th International Conference on Tissue Engineering (Kos, 2014)
where she also received the Aegean Conference Travel Award in recognition of her
performance in the rapid-fire session.
178
“Success is not final, failure is not fatal:
it is the courage to continue that counts.”
- Winston Churchill
179