There are two types of bone tissue: compact and spongy. The names imply that the two types of differ in density, or how tightly the tissue is packed together. There are three types of cells that contribute to bone homeostasis. Osteoblasts are bone-forming cell, osteoclasts resorb or break down bone, and osteocytes are mature bone cells. An equilibrium between osteoblasts and osteoclasts maintains bone tissue.¹

Compact Bone

Compact bone consists of closely packed osteons or haversian systems. The osteon consists of a central canal called the osteonic (haversian) canal, which is surrounded by concentric rings (lamellae) of matrix. Between the rings of matrix, the bone cells (osteocytes) are located in spaces called lacunae. Small channels (canaliculi) radiate from the lacunae to the osteonic (haversian) canal to provide passageways through the hard matrix. In compact bone, the haversian systems are packed tightly together to form what appears to be a solid mass. The osteonic canals contain blood vessels that are parallel to the long axis of the bone. These blood vessels interconnect, by way of perforating canals, with vessels on the surface of the bone.¹

Spongy (Cancellous) Bone

Spongy (cancellous) bone is lighter and less dense than compact bone. Spongy bone consists of plates (trabeculae) and bars of bone adjacent to small, irregular cavities that contain red bone marrow. The canaliculi connect to the adjacent cavities, instead of a central haversian canal, to receive their blood supply. It may appear that the trabeculae are arranged in a haphazard manner, but they are organized to provide maximum strength similar to braces that are used to support a building. The trabeculae of spongy bone follow the lines of stress and can realign if the direction of stress changes.¹

Figure1 Compact bone and Spongy Bone¹

2. BONE HEALING

The soft-tissue healing of a wound creates a fibrous scar but bone is unique in its scarless regenerative capacityFollowing damage to the musculoskeletal system, disruption of blood vessels leads to activation of the coagulation cascade and formation of a hematoma, which encloses the fracture area (Figure 2a). Removal of the hematoma significantly attenuates repair, and transplantation of the hematoma produces new bone,

consistent with the angiogenic activity of the hematoma. Inflammatory cells, fibroblasts, and stem cells are recruited to the site, and new blood vessels are formed from pre-existing ones (i.e. angiogenesis). The inflammatory response is associated with pain, heat, swelling, and the release of several growth factors and cytokines that have important roles in repair. Initially, granulation tissue forms at the ends of bones, gradually being replaced by fibrocartilage, in a manner seemingly related to the vascular pattern. Meanwhile, the periosteum undergoes direct bone formation, or intramembranous ossification, to create an external callus (Figure 2b). Subsequently, the internal callus becomes mineralized with calcium hydroxyapatite, to form a hard callus of woven bone (Figure 2c). In the final, remodeling large fracture callus is replaced by secondary lamellar bone; the size of the callus is reduced to that of pre-existing bone at the damage site, and the vascular supply reverts to a normal state (Figure 2d).²

Figure 2 The stages of fracture repair.²

3. TISSUE ENGINEERING

Tissue engineering has been defined as the application of scientific principles to the design, construction, modification and growth of living tissues using biomaterials, cells and factors, alone or in combination.³

Recently, tissue engineering, which applies methods from engineering and life sciences to create artificial constructs to direct tissue regeneration, has attracted many scientists and surgeons with a hope to treat patients in a minimally invasive and less painful way. The tissue engineering paradigm is to isolate specific cells through a small biopsy from a patient, to grow them on a three-dimensional scaffold under precisely controlled culture conditions, to deliver the construct to the desired site in the patient’s body, and to direct new tissue formation into the scaffold that can be degraded over time.⁴

Figure 2 Tissue engineering strategy⁵

Key factors involved in tissue engineering

Identification of a suitable cell type

The main components required for tissue engineering are cells, scaffolds and growth factors. One major problem is the identification of a cell type that is both clinically viable

and suitable for tissue engineering. Mesenchymal stem cells have an inherent potential to differentiate into osteogeneic,chondrogeneic, adipogeneic and myocardiac cell lineages.

the two requirements for tissue regeneration and organ substitution are an increase in cell proliferation and the maintenance of their biological functions. Research into the identification of cell growth and differentiation factors, and either the production of these factors on a large scale or improvement of the chemical components of cell culture media and substrates is still necessary.⁶

Scaffold development

The ECM that surrounds cells in the body not only physically supports cells but also regulates their proliferation, differentiation and morphogenesis18. Such a scaffold, therefore, needs to be developed for in vitro tissue reconstruction as well as for cell-mediated tissue regeneration in vivo.⁵

Ideally, a scalfold should have the following characteristics⁷:

(i) three-dimensional and highly porous with an interconnected pore network for cell growth and flow transport of nutrients and metabolic waste;

(ii) biocompatible and bioresorbable with a controllable degradation and resorption rate to match cell/tissue growth in vitro and/or in vivo;

(iii) suitable surface chemistry for cell attachment, proliferation, and differentation and

(iv) mechanical properties to match those of the tissues at the site of implantation

Growth factors

Successful tissue regeneration cannot always be achieved by the combination of cells and their scaffold alone. In such cases, a suitable growth factor to promote tissue regeneration is required.⁶

4. BONE TISSUE ENGINEERING

A large number of bone fractures have been treated by bone grafting.. With an increase of the mean population age, the development and optimization of bone regeneration techniques represents a major clinical need for many countries. Autografts have limitations due to the necessity of an additional surgery, limited donor bone supply, anatomical and structural problems and inadequate resorption rate during healing. Allografts have the disadvantage of a potential immune response, transmitting diseases, and they may induce the loss of osteoinduction. Metals alone or coated with bioactive

and bioinert ceramics have been used for load-bearing orthopedic applications, but problems due to metals corrosion, ceramics–metal interface wear, and dense fibrous tissue formation on the bone–implant interface may occur.⁸

Bone tissue engineering is a new research area with clinical applications in bone replacement on orthopedic defects, bone neoplasia and tumors, pseudoarthrosis treatment, stabilization of spinal segments, as well as in maxillofacial, craniofacial, orthopedic, reconstructive, trauma and neck and head surgery. It may provide solutions for generating a new bone tissue with good functional and mechanical qualities, reducing the risks and expenses of using autografts, allografts and metals.⁸

The research program for tissue engineering bone is classified into six phase⁷

1) Fabrication of bioresorbable scaffold

2) Seeding of the osteoblasts/chondrocytes populations into the

polymeric sca!old in a static culture (petri dish)

3) Growth of premature tissue in a dynamic environment (spinner flask)

4) Growth of mature tissue in a physiologic environment (bioreactor)

5) Surgical transplantation

6) Tissue-engineered transplant assimilation/remodeling

5. APPLICATIONS OF BIOMATERIALS FOR BONE TISSUE ENGINEERING

Biomimetic materials

The development of biomaterials for tissue engineering applications has recently focused on the design of biomimetic materials that are capable of eliciting specific cellular responses and directing new tissue formation mediated by biomolecular recognition, which can be manipulated by altering design parameters of the material. The surface modification of biomaterials with bioactive molecules is a simple way to make biomimetic materials. The early work has used long chains of ECM proteins such as fibronectin (FN), vitronectin (VN), and laminin (LN) for surface modification. Biomaterials can be coated with these proteins, which usually have promoted cell adhesion and proliferation. Since the finding of the presence of signaling domains that are composed of several amino acids along the long chain of ECM proteins and primarily interact with cell membrane receptors, the short peptide fragments have been used for surface modification in numerous studies. the short peptide sequences are relatively more stable during the modification process than long chain proteins such that nearly all modified peptides are available for cell binding. In addition, short peptide sequences can be massively synthesized in laboratories more economically. The biomimetic material modified with these bioactive molecules can be used as a tissue engineering scaffold that potentially serves as an artificial ECM providing suitable biological cues to guide new tissue formation.⁴

The most commonly used peptide for surface modification is Arg-Gly-Asp (RGD), the signaling domain derived from fibronectin (FN) and laminin (LN). Additionally, other peptide sequences such as Tyr-Ile-Gly-Ser-Arg (YIGSR), Arg-Glu-Asp-Val (REDV), and Ile-Lys-Val-Ala-Val (IKVAV) have been immobilized on various model substrates. A number of materials including glass, quartz, metal oxide, and polymers have been modified with these peptides and characterized for cellular interaction with their surfaces.⁴

It has been found that many glycoproteins in the ECM are expressed during new bone formation. For example, a variety of matrix proteins such as osteopontin, thrombospondin, and bone sialoprotein, which were important in bone cell migration, proliferation, matrix deposition, and mineralization, have been identified. The integration of biomimetic scaffolds into bone tissue takes place at the material–tissue interface that is often determined by initial cell and substrate interactions. Thus, it may be necessary to design biomimetic scaffolds modified with bone matrix proteins that serve as biological cues for cell–matrix interactions to promote bone growth.⁴

Poly(lactic acid) (PLA) and poly(lactic-co-glycolic acid) (PLGA) were coupled with RGD peptides using poly(llysine) and fabricated as scaffolds for the growth of osteoblasts. Peptide modified biodegradable polymers demonstrated enhanced initial cell attachment and significant expression of osteogenic phenotype.⁴

In addition to RGD peptide sequences, other adhesive peptides that can be recognized by polysaccharide molecules in the cell membrane were used for the design of biomimetic materials for bone tissue engineering. Particularly, the heparin binding domain in FN has been investigated as a candidate to mediate specific binding with bone derived cells. Dalton et al. exhibited that monoclonal antibody Mab32, which blocks a heparin binding domain, partially inhibited osteoblast adhesion to the heparin binding domain treated surface. These results suggest that heparin binding domains as well as other cell adhesive sequences are actively involved in specific interaction with osteogenic cells. ⁴

silk fibroin net

Silk fibers have unique properties that fulfill many of the requirements for biomaterial scaffolds. Silk exhibits high strength and flexibility and permeability to water and oxygen. In addition, silk can be made into fibers, sponges or membranes. These characteristics make silk an excellent substrate for biomedical applications such as implant biomaterials, cell culture scaffolds, cell carriers or biological fluid filtering systems. Undesirable immunological problems attributed to the sericin protein of silk limited the use of silk. Purified silk fibroin, which remains after removal of sericin, exhibits low immunogenicity and good blood compatibility ⁹.

Figure 4. On the left, Scanning electron images of non-woven silk fibroin nets are seen (low-magnification image showing individual fibers and branching) On the right, Confocal images of calcein-AM stained osteoblast cells grown 3–7 days on non-woven silk fibroin nets are seen.⁹

It has been shown that silk fibroin nets support the attachment, spread, and growth of osteoblast The extensive growth of osteoblast on the nets point to the possible use of silk fibroin as a biomaterial for bone regeneration.

Biphasic calcium phosphate ceramics

Bioactive ceramics, namely hydroxyapatites (HA), hydroxyapatite/tricalcium phoshatates (HA/TCP) and bioactive glasses, have been used as scaffolds for bone reconstruction for many years. They are termed “bioactive” because they form an interfacial bond with tissues upon implantation and bone tissue formation as result of surface modification when exposed to interstitial fluids. However, compared with human cortical bone, they have lower fracture toughness and higher elastic moduli. Therefore, it is desirable to develop bioactive materials with improved mechanical properties.¹⁰

It has been demonstrated that MSCs can induce bone repair when implanted in vivo in combination with biphasic calcium phoshate, specifically hydroxyapatite/tricalcium phoshate. HA/TCP implants were loaded with human MSCs cells. Harlan Nude rats were subjected to the femoral gap. Implants were placed, with one limb receiving a MSCs loaded implants and contralateral limb receiving a cell-free scaffold. Radiografic, histological, and mechanical analysis was used to evaluate the regenerative capacity of the human MSCs. Ostogenesis was apparent at 4 weeks following implantation in the human MSC-loaded implants, but absent in the cell-free implants. After 8 weeks, the MSC-loaded implants were well integrated with the host. The result of mechanical testing showed that limbs treated with human MSC-loaded implants has statistically greater strength and stiffness values as compared to limbs treated with the scaffold alone.¹⁰

Figure 5 Radiographs of rat femoral defects (a) prior to implantation, (b) immediately post-implantation, (c) treated with HA/TCP loaded human MSCs at 12 weeks after implantation (d) treated with HA/TCP alone at 12 weeks after implantation¹⁰

Hydroxyapatite–chitin material

Chitin, the precursor of chitosan, is one of the most abundant natural polymers. Chitin has been shown to enhance biological self-defense function in animals, accelerate wound healing, and be used as a biocompatible and absorbable material in both animal and plant tissues. Calcium hydroxyapatite (HA) is the main component of teeth and bones in vertebrates. In this study, HA–chitin materials were prepared and evaluated for general biocompatibility, and subsequently seeded with mesenchymal stem cells harvested from NZW rabbits and evaluated for their performance as potential bone substitutes.¹¹

Preliminary evaluation

Cellular studies

MTT test results showes that HA–chitin materials were non-cytotoxic to the mouse fibroblast, human fibroblast or human bone cells¹¹.

Direct contact tests give an idea of the response of a cell layer to the contact with the HA–chitin materials. After 3 days of culture, the osteoblasts started to adhere and attach as isolated cells onto the surface of the HA–chitin thin-films. After 14 days of culture, cells had started to proliferate on the surface of the matrixes, therefore forming clusters of cells as shown in Fig. .These results indicated that the support provided by HA–chitin did not alter the biological activity of the osteoblasts as they grew normally and retained their characteristics.¹¹

Figure 6 SEM photomicrograph of osteoblasts proliferating on the surface of a 25% HA–chitin thin-film after 14 days of culture.¹¹

intramuscular implantation

Sprague-Dawley rats were used in intramuscular implantation test. They were anesthetized The implants were surgically implanted bilaterally into the dorsal-lumbar musculature.

The results after 14 days of intramuscular implantation:

ü HA–chitin thin-film was developed to promote calcification,

ü HA–chitin was surrounded by a dense and thick fibrous capsule

Figure7 Morphology of the muscle tissue around HA–chitin implants after 14 days of implantation

(A) 25% HA–chitin sample showing normal fibrous connective tissue (G) and newly formed capillaries filled with red blood cells (B). Calcification (C) is observed in the

middle of the implant (I).

(B) Interface between a 50% HA–chitin sample (top) and the scar tissue (bottom).¹¹

The results after 3 months of intramuscular implantation:

ü HA–chitin matrixes were degraded faster than HA–chitin thin-film

ü the thickness of the fibrous capsule around the HA–chitin materials had decreased (Fig.C), due to the organization and the maturation of the fibrous scarring process.

Figure 8 after 3 months implantation

(D) Resorption of a 50% HA–chitin matrix after 3 months of implantation. The material was invaded and divided by the granulation tissue(G), leading to the production of remnants (R) of various shapes detached from the main body of the implant (I)¹¹

It is observed that degradation of chitin-HA material in vivo was rapid.This may be disadvantages for the application of HA-chitin materials to load bearing situations since they may not maintain strong mechanical support against rapid degradation.¹¹

Induction of osteblasts from mesenchymal stem cells

The mature osteoblast phenotype is characterized by their ability to synthesize membrane associated alkaline phosphatase, bone matrix molecules including collagen type I and non-collagenous proteins such as osteocalcin, bone sialoprotein and osteopontin, proteoglycans, hormonal and other growth factor receptors. Osteocalcin is widely recognized as a bone-specific protein or an exclusive product of the osteoblast/osteocyte.¹¹

In this work, alkaline phosphatase (AP) and osteocalcin were used as indicative markers for osteoblasts. The MSCs were first induced to osteoblasts with dexamethasone in vitro. The induced osteoblasts showed typical cuboidal and basophilic appearance¹¹

Figure 9 The induced osteoblasts showed typical cuboidal and basophilic appearance under Masson stain¹¹

In vivo model studies

Both cell loaded and cell-free matrixes were implanted into rabbit femur. The results from the 2 months’ implantation suggest that both cell-free and cell-loaded porous HA–chitin matrixes promoted the in-growth of surrounding tissues, with the cell-loaded HA–chitin matrixes being the better performer. MSC induced osteoblast loaded matrixes showed that the osteoblast did not only survive and migrate when implanted in vivo, but continually proliferated in the matrixes or degraded matrixes.¹¹

6. CONCLUSION

The developing appropriate scaffold material for bone tissue engineering is seem critically important. Scaffold has to be non-cytotoxic, nonallergic and nonimmunogenic. It has to provide proper surface chemistry and surface microstructure for optimal cell-substrate interaction and acts as a template to direct and organize cell growth, ECM formation and vascularization. Another important point is that degradation rate of scaffold should match the degradation rate of tissue regeneration. If it is degraded faster as in the cases of HA–chitin materials, it will not support tissue mechanically.

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