Index of /rozprawy2/11202

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Wydział Inżynierii Materiałowej i Ceramiki

Katedra Biomateriałów

Dissertation for the degree of Doctor of Philosophy

HIGHLY POROUS CERAMIC SCAFFOLDS FOR BONE

TISSUE ENGINEERING

mgr inż. Łucja Rumian

Scientific Supervisor: prof. dr hab. inż. Elżbieta Pamuła

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“Every science comes with its own pseudo-science, a bizarre distortion that comes from a certain kind of mind.”

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ACKNOWLEDGEMENTS

Firstly, I would like to express my sincere gratitude to my supervisor prof. dr hab. inż. Elżbieta Pamuła for her guidance and mentorship through all those years, the help in completing my Ph.D. thesis and all good advice. My sincere thanks also goes to dr inż. Urszula Cibor for her good advice in the area of microencapsulation and emulsions. I am particularly indebted to dr inż. Małgorzata Krok-Borkowicz, for her help in cell cultures. I thank my fellow members of the Pamuła group for the stimulating discussions and lots of fun we had together: Iwona Wojak-Ćwik, Krzysztof Pietryga, Katarzyna Reczyńska and Bartosz Mielan. I also want to thank dr hab. Monika Brzychczy-Włoch for collaboration at microbiological experiments.

This work would not have been possible without the support from Håvard J. Haugen, professor of the Department of Biomaterials at the University of Oslo. The discussions I had with him were invaluable. I would like to gratefully acknowledge my Oslo supervisor, Hanna Tiainen, PhD, for introducing me to the world of ceramics and sintering. I would like to thank Jonas Wengenroth for his help with micro-computed tomography. Special thanks to Grażyna Jonski, Shahbaz Yousefi and Sonny Langseth-Manrique for their patients and invaluable help during my stays in Oslo.

As regards the osteogenesis experiment, I thank Dieter Scharnweber, professor of the Max Bergann Center at the TU Dresden, for his scientific and financial support. I would like to extend my gratitude to Dr. rer. nat. Cornelia Wolf-Brandstetter, my supervisor in Dresden. Thanks for graciously lending your time and expertise in PBMCs culture lab work. Special thanks to Dipl.-Ing. (FH) Heike Zimmermann, for her infinitive patience, warmth and invaluable help.

I am indebted to all my friends who supported me during those years. My final words go to my family. I want to thank my family, whose love and guidance is with me in whatever I pursue.

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CONTRIBUTORS AND FUNDING SOURCES

This study was financed from Polish National Science Centre (Grant no. 2013/09/N/ST8/00309) and Norwegian Research Council grant no. 228415. The research in Dresden was financed by the binational Research Development and Training Network "German Polish Network for Bone Engineering" (GO BONE), financially supported by the German Federal Ministry of Education and Research (BMBF).

Sodium alendronate used in the experiments was a gift from Polpharma S.A., Poland, whereas salmon calcitonin was graciously given by Jelfa S.A., Poland.

Copolymer PLGA was kindly provided by prof. dr. hab. Piotr Dobrzyński from Centre of Polymer and Carbon Materials of the Polish Academy of Sciences in Zabrze.

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LIST OF ABBREVIATIONS

%EE encapsulation efficiency

%LE loading efficiency

µCT micro-computed tomography

Alg sodium alginate

Aln sodium alendronate

BMPs bone morphogenetic proteins

BSA bovine serum albumin

CaCl2 calcium chloride

Calc salmon calcitonin

CaP calcium phosphate

DBM demineralised bone matrix

DCM dichloromethane

DO distraction osteogenesis

EDTA ethylenediaminetetraacetate

EMEM Eagle’s minimal essential medium

FBS foetal bovine serum

Gent gentamicin sulphate

HAp hydroxyapatite

IL-1ß interleukin 1 beta

IlL-6 interleukin 6

MCSF macrophage colony-stimulating factor

MPs microparticles

MRSA methicillin-resistant Staphylococcus aureus MSSA methicillin-susceptible Staphylococcus aureus

nHAp nanocrystalline hydroxyapatite

OPA orthophthaldialdehyde

OSS one-step sintering

PBS phosphate buffered saline

PCL poly(ε-caprolactone)

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PLGA poly(L-lactide-co-glycolide)

PMMA poly(methyl methacrylate)

PRP platelet-rich plasma

PTH parathyroid hormone

PTH 1-34 teriparatide

PVA polyvinyl alcohol

RANKL receptor activator of nuclear factor kappa-B ligand

S. aureus Staphylococcus aureus

S. epidermidis Staphylococcus epidermidis

SEM scanning electron microscope

Staphylococcus spp. Staphylococcus species

TCP tricalcium phosphate

TCPS tissue culture polystyrene

TiO2 titanium dioxide

TNF-α tumour necrosis factor alpha TRAP tartrate-resistant acid phosphatase

TSS two-step sintering

UHQ-water ultra high quality water

Vanc vancomycin hydrochloride

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Chapter 1. GENERAL INTRODUCTION

1.1 Outline and aims of the thesis

Bone is the main supporting system in human body. Organic and mineral components of bone provide excellent tensile and loading strength. Bone plays key roles in human physiology such as protection, movement, support of organs, blood production, mineral storage and homeostasis, stem cells housing, blood pH regulation, and others. Bone possesses the intrinsic capacity for regeneration as part of the repair process in response to injury, as well as during skeletal development or continuous remodelling throughout adult life [1]. As such, the process of fracture healing recapitulates bone development and can be considered a form of tissue regeneration. Bone regeneration is comprised of a well-orchestrated series of biological events of bone induction and conduction. Those involve a number of cell types and intracellular and extracellular molecular signalling pathways, with a definable temporal and spatial sequence, in an effort to optimise skeletal repair and restore skeletal function [2]. When this complicated system ceases to work properly the quality of life of the individual degrades rapidly leading even to premature death. Metabolic diseases, bacterial infections along with traumatic injuries, orthopaedic surgeries and primary tumour resections lead to or induce bone defects. That, in the ageing societies, becomes a serious burden both socially and economically [3].

During the planning stage of the treatment for bone loss, several factors must be taken into account: the quality of the soft tissue envelope, the quality of vascular supply and the presence or absence of an infection [4]. Bone loss in certain anatomical locations has a more favourable prognosis due to better blood supply and corresponding osteogenic potential. In standard treatment protocols for bone defects the positive outcome rates vary significantly depending on affected site or overall patient’s health condition. The age of the patient, the presence of chronic disease (e.g. diabetes mellitus), use of medications, alcohol consumption and tobacco usage may hamper bone regenerative abilities [5]. Present therapeutic approaches include bone graft transplants (autologous or allogeneic grafts), implants of different biomaterials and distraction osteogenesis (Ilizarov technique), but none of them can be described as fully satisfactory.

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For all above-mentioned reasons the focus of this thesis was to develop viable porous ceramic scaffolds as bone graft substitutes which at the same time will address the problem of common causes of the bone loss, such as microbial infections or poor condition of bone tissue resulting from metabolic diseases (such as osteoporosis) or bone tumours or bone metastasis.

Chapter 1 contains a systemic review of bone defects, their main causes, classification followed by brief summary of modern methods of bone loss treatment. A special focus is given to bone tissue engineering approach, what are the main advantages and disadvantages in comparison to conventional treatment methods, which are the most active areas of current research addressing those problems.

Chapter 2 of this thesis focuses on obtaining titanium dioxide (TiO2) scaffolds

with optimal mechanical properties, microstructure and architecture. Ceramic TiO2 has

been shown to have superb biocompatibility in contact with bone tissue. It also promotes osseointegration. Scaffolds were fabricated by polymer sponge replication method. Proposed approach should combine optimal manufacturing method, appropriate architecture of the pores in the scaffolds, and biocompatibility of TiO2 to obtain

load-bearing constructs for bone tissue engineering.

In Chapter 3 the main aim was to improve manufacturing process of polymeric poly(L-lactide-co-glycolide) (PLGA) microparticles (MPs) loaded with four different drugs: gentamicin sulphate (Gent), vancomycin hydrochloride (Vanc), sodium alendronate (Aln) and salmon calcitonin (Calc). Parameters such as encapsulation efficiency, loading efficiency and particle size distribution were main indicators in fabrication process.

Post-surgical infections stand at the forefront of concern to the modern implantology. Therefore a viable systems ensuring constant local antibiotic delivery are in high demand. Chapter 4 covers my proposal for dealing with this problem by combining biocompatibility and osteoconductivity of TiO2 scaffolds (Chapter 2) with controlled Gent

and Vanc release from PLGA MPs described in Chapter 3. The main goal was to immobilize within the scaffolds Gent- or Vanc-loaded MPs while maintaining biocompatibility of the scaffolds, as well as an appropriate antibiotic release kinetics from MPs. To confirm utility of the systems drug release, cytocompatibility in contact with osteoblast-like cells and antimicrobial activity on Staphylococcus spp., both commercial and clinical strains isolated from infected bones and joints were tested.

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3 Chapter 5 covers a slightly different approach to the bone remodelling after implantation, by addition of Aln (the main representative of bisphosphonates) or Calc to diminish osteoclasts’ activity at the implantation site. Similarly as in Chapter 4 the MPs on the scaffolds surface and pore walls to assure a suitable release dosage of the drugs were immobilised. Drug release kinetics was tested, followed by cytocompatibility assessment in contact with osteoblast-like cells and the influence of the systems on peripheral blood mononuclear cells differentiation towards osteoclasts.

1.2 Bone defects and their causes

Bone loss may be associated with either bone defect or structural loss within bone tissue. Structural loss may be caused by osteopenia, osteoporosis or other metabolic diseases. Primary causes for bone defects are open fractures and defects, gunshot wounds or osteoclastic tumours. Secondary bone loss may be the result of tumour resection, post-surgical infection or non-unions [4,6].

Critical size defect in bone is “an orthotopic defect that will not heal without intervention” [7]. Bone tissue has the intrinsic capacity for regeneration, which allows the self-repair of small bone lesions. However, in the case of critical size defects or the imbalance in local environment, an adequate remodelling of the damaged or missing tissue is impossible. Provided that the patient suffers also some kind of metabolic disorder (e.g. osteoporosis) the healing of the defect poses an even bigger challenge. The defects can be considered critical in relation to the skeletal segment involved and the length of bone loss: 3 cm for the forearm, 5 cm in the femur and tibia, 6 cm in the humerus [8].

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Figure 1-1 Stages of normal bone healing in non-stabilized fracture: (A) inflammation, (B) soft callus and (C) hard callus phase [9].

The bone repair process (Figure 1-1) starts with anabolic phase and is initiated by the inflammatory response and formation of fracture hematoma. An increase in tissue volume is related to the recruitment and differentiation of stem cells that form skeletal and vascular tissues. A cartilaginous callus forms directly adjacent to the fracture line. At the edges of the newly formed cartilage tissue, the periosteum swells and primary bone formation is initiated. Ingrowth of capillaries into the callus and increased vascularity follows. Closer to the fracture gap, mesenchymal progenitor cells proliferate and migrate through the callus, differentiating into fibroblasts or chondrocytes, each producing their characteristic extracellular matrix and slowly replacing the hematoma. Catabolic activity, such as cartilage resorption, predominates in second stage of fracture healing, specific

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5 anabolic processes continue to take place; secondary bone formation is initiated as the cartilage is resorbed and primary angiogenesis continues as the nascent bone tissues replace the cartilage. Subsequently, when bone remodelling begins, the first mineralized matrix produced during primary bone formation is resorbed by osteoclasts, and then the secondary bone laid down during the period of cartilage resorption is also resorbed. As the bony callus tissue continues to be resorbed, this prolonged period is characterized by coupled cycles of osteoblast and osteoclast activity in which the callus tissues are remodelled to the bone’s original cortical structure. The fracture repair is essentially dependent on adequate stability, formation of new vasculature and remodelling of the callus by osteoclasts [2,9,10]. In the following paragraphs three main causes of the bone defects formation will be discussed in detail: trauma, osteomyelitis and osteoporosis.

1.2.1 Trauma and post-implantation related bone defects

Trauma related bone loss may occur either during extrusion of fragments at the time of injury or during debridement when impaired bone is removed, which creates a defect. In Europe predominant trauma type is blunt force trauma. Therefore most of the skeletal defects are created at the time of debridement. In the areas where gunshot wounds are more common, extrusion of bone fragments can occur at the time of injury [5]. In contrast to blunt force trauma, ballistic injuries result in significant bone and soft tissue loss, whose severity is not always apparent at initial presentation. Reconstruction of these defects is often complicated by tissue ischaemia, necrosis and infections [11]. Ten-year data from a fracture registry indicated that 0.4% of all fractures that were treated at a Level-I orthopaedic trauma unit were associated with bone loss [5]. The decision to salvage the severely injured limb remains objective based on muscle and joint damage as well as the potential for neurological and vascular recovery. Of particular value in high-energy trauma is the ability to deal with bone, joint and soft tissue injury with minimal donor site morbidity [12]. Musculoskeletal injuries are a substantial burden in developing countries such as India; the problem is complex, multidimensional, and can only be solved through a multidisciplinary, multi-sectoral effort. In highly developed countries, such as in the United States, it is estimated that delayed or impaired healing will occur in 5–10% of the 5.6 million fractures that occur annually, and up to 10% of all fractures will require additional surgical procedures for impaired healing [13].

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Long-bone defects are one of the most challenging problems associated with limb reconstruction following high-energy extremity trauma. A common result of lack of bone integrity reconstitution is amputation [14]. The most common site of bone loss after fracture is the tibia, because of its subcutaneous position which predisposes it to open fracture and extrusion of bone. Open fractures of the upper limb and axial skeleton are less common and bone loss is seldom encountered in these locations [5].

Aseptic loosening is the most common cause of failure of cemented and uncemented joint replacements. Aseptic loosening can be the result of inadequate initial fixation, mechanical loss of fixation over time, or biologic loss of fixation caused by particle-induced osteolysis around the implant. The causes of particle accumulation vary from implant interface wear, micro-motion occurring in response to corrosion, oxidative reactions, and minor pathogen contaminations. Loosening of implant components is thought to occur in over 10% of cases within 20 years of a primary hip arthroplasty [15,16]. Bone resorbing osteoclasts develop from the monocyte/macrophage lineage under the influence of macrophage colony-stimulating factor (MCSF) and receptor activator of nuclear factor kappa-B ligand (RANKL). They remove old, injured and “unneeded” bone and at the same time stimulate bone regeneration via several mechanisms that couple bone resorption to bone formation [17]. Biomaterial particle debris elicits a heavy macrophage response which has been associated with aseptic loosening in joint replacement, and it has been suggested that these macrophages actually resorb bone [18]. Studies in vivo and in

vitro have demonstrated that monocytes and macrophages engulf particles which in turn

induce secretion of tumour necrosis factor alpha (TNF-α), interleukin 1 beta (IL-1ß), interleukin 6 (IL-6) and prostaglandin E2 (PGE-2) that stimulate differentiation of osteoclast precursors into mature osteoclasts. The accumulation of macrophages is important as osteoclasts are derived from this lineage of cells [19]. This effect leads to bone resorption and in excess can induce implant failure. While in debate whether osteoclasts contain foreign material in their cytoplasm, it also has been shown, that macrophages, which have phagocytosed particles, are capable of osteoclast differentiation [20].

Long-term stability of dental implants demands its osseointegration in the alveolar bone. The osteoblast lineage cells, following their recruitment and differentiation, synthesize the bone collagenous organic matrix and control its mineralization. A proper remodelling process involving the coupled activities of the osteoblasts and the osteoclasts plays a key role in homeostasis and mechanical performance of the peri-implant bone.

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Peri-7 implantitis is characterized as an inflammatory reaction that affects the hard and soft tissue, which results in loss of supporting bone and pocket formation surrounding the functioning osseointegrated implant [21]. The following factors are associated with the occurrence of peri-implantitis: imbalanced occlusal force, smoking habit, faults in material and techniques, poor bone quality at the implant area, poor oral hygiene, bacterial infection, diabetes, the particles released from implant, etc. [22,23]. In the case of peri-implantitis an immune-inflammatory induced osteoclastogenesis is its central pathological hallmark. The underlying mechanism of the inflammatory process leading to bone resorption appears to be centred on RANKL. As mentioned above, cells involved in the local host reaction produce several cytokines that induce the synthesis of this osteoclastogenic factor by stromal cells and osteoblasts. Additionally, activated immune cells are also RANKL producers [22].

1.2.2 Osteomyelitis

Osteomyelitis is a severe, progressive inflammatory bone disease caused by microbial infection or auto-inflammatory processes. The associated inflammation involves the bone marrow and the surrounding tissues. Traditionally osteomyelitis has been classified, based on the source of the infection, as haematogenous, or secondary to a contiguous focus of infection, or those associated with peripheral vascular disease [24]. The first one, is bone infection that has been seeded through the bloodstream and it is primarily a disease of children. The second one is seen most often after trauma or surgery, and is caused by bacteria which gain access to bone by direct inoculation (e.g. a contaminated compound fracture) or extension to bone from adjacent contaminated soft tissue (for example, a prosthetic joint contaminated at the time of implantation) [25,26].

A second classification, which applies irrespective of the underlying source of infection, is based on the time between the onset of symptoms and diagnosis, and makes a distinction between acute (<2 weeks), subacute (2 weeks – 3 months) and chronic infection (>3 months) [27]. During acute phase of osteomyelitis an acute suppurative inflammation occurs, in which bacteria or other microorganisms are rooted. This leads to severe immunological response of the organism which contributes to tissue necrosis and destruction of bone trabeculae and bone matrix. In the next stage, vascular channels are destroyed by the inflammatory process, which may lead to ischemic necrosis. This dead bone, known as a sequestrum, can act as a non-living surface for biofilm attachment, which

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continue to harbour bacteria despite antibiotic treatment [25,28]. Chronic osteomyelitis is commonly characterised by the persistence of microorganisms, low-grade inflammation, and the presence of dead bone and fistulous tracts [28]. The hallmark of chronic osteomyelitis is the simultaneous presence of microorganisms and necrotic bone [24].

Bacteria of the genus Staphylococcus are the principal causative agents of two major types of infection affecting bone – septic arthritis and osteomyelitis, which involve the inflammatory destruction of joint and bone [29]. Staphylococcus aureus is by far the most frequent isolated causative agent of infection-induced chronic osteomyelitis [24,30]. The ability of S. aureus to adhere is thought to be crucial for the early colonisation of host tissue, implanted biomaterials, or both. It can also survive in osteoblasts after internalization which protect it from immune cells but also from the antibiotics that may not penetrate the cells [31]. Another bacteria from Staphylococci family is Staphylococcus

epidermidis. Normally considered as a harmless human skin colonizer, yet causes

approximately 20% of prosthetic joint infections and fracture fixation infections [32]. In general Staphylococcus spp. can produce a multi-layered biofilm embedded within a glycocalyx, or slime layer. The presence of implants is a predisposing factor in the development of infection since on the surface of the implant a protein coating is forming. That coating is an excellent source of attachment for any bacteria introduced or remaining after debridement surgery [33]. A number of other bacterial pathogens cause osteomyelitis, such as Streptococcus pyogenes, Acinetobacter baumannii, Pseudomonas aeruginosa, and

Escherichia coli [26,29,30,34]. The patients prone to osteomyelitis are those who

experience a higher frequency of disorders that lead to infection, such as orthopaedic surgeries and diabetes mellitus [34].

Osteomyelitis is usually associated with open fracture surgery, bone reconstruction surgery, or orthopaedic implants. Post-traumatic osteomyelitis is currently one of the greatest challenges in orthopaedic surgery. Infection rates after orthopaedic surgery range from 0.4% to 6%, depending on the type of the surgery that is done [35]. Infection of the calvarium after calvarial reconstructions occurs in 2.5 up to 6.5% of cases [36]. Vertebral osteomyelitis occurs in 2 – 7% of patients. Periprosthetic infection is the leading cause of revision total knee arthroplasty and the third most common cause for total hip arthroplasty; the incidence of these periprosthetic infections has been estimated at 1 to 3% [37,38]. In adult patients, it is estimated that 47 – 50% of all osteomyelitis are post-traumatic [39]. Rubin et al. calculated the direct medical costs for treatment of a patient

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9 with a bony osteomyelitis to be $35,000 [40]. Moreover, several studies have shown that treatment costs of osteomyelitis may well rise up to 500,000.00 € per case [41].

1.2.3 Osteoporotic fractures

Osteoporosis is a systemic skeletal disorder characterized by bone loss and microarchitectural deterioration of the trabecular network. The increased bone turnover reduces the mean degree of mineralization of bone. Up to 9 million people worldwide every year are affected by fragility fractures, which are common in the elderly population [42,43]. Globally, the number of aged in the population is expected to double by 2040; consequently the number of osteoporotic fractures will increase greatly. There is a huge cost associated with osteoporosis in terms of morbidity, mortality, and the financial impact on society. Osteoporotic fractures cost the U.S. around $17.9 billion per year, the UK around £1.7 billion per year. The worldwide economic cost of osteoporosis in 1998 was $34.8 billion and is expected to rise to $131.5 billion by 2050 [44].

Patients with osteoporosis suffer a reduction in bone mineral density that leads to an increased risk of fracture. The bone deposition by osteoblasts is in imbalance with osteoclastic bone resorption. That leads to a net loss of bone over time. Therapeutic agents for osteoporosis could increase bone strength by three different but interrelated effects on bone tissue: prevention of bone loss, an increase in the volume of bone matrix and an increase in the degree of mineralization [45]. Thus, osteoporotic patients are commonly treated with anabolic agents for stimulation of bone formation (e.g., parathyroid hormone (PTH)) or anti-resorptive agents (e.g., bisphosphonates, Calc, raloxifene and estrogen) which act by inhibiting bone resorption [46,47].

The most common fractures connected to osteoporosis are hip fractures, vertebral fractures and distal forearm fractures [44,48,49]. This is connected with high proportion of cancellous bone in those regions. The most devastating complication of osteoporosis is a hip fracture [42]. 10 – 20% more women die than expected for age within the first year after fracture occurrence, and the excess mortality is even greater for men [50]. Hip fracture rates are significantly higher among women, the differences in incidence between age groups are much larger in women than in men [44]. This is due to rapid drop in estrogen levels after menopause beginning, which causes 1 to 3% drop per year in bone mineral density for as long as ten years [51].

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1.3 Methods of bone defects treatment

1.3.1 Current clinical treatments for bone repair and regeneration

Significant bone defect or post-traumatic complications may require bone grafting in order to fill in the defect. Bone grafting is a common surgical procedure, more than 500,000 procedures are happening annually in the U.S. and 2.2 million worldwide in order to repair bone defects in orthopaedics, neurosurgery and dentistry [52]. Osteogenesis, osteoinduction and osteoconduction are the three essential elements of bone regeneration along with the final bonding between host bone and grafting material which is called osseointegration. Osteoinductivity is the ability of a graft to actively stimulate or promote bone formation. Osteoinduction implies the recruitment of immature cells and the stimulation of these cells to develop into preosteoblasts. Osteoconductivity, a very important property of a bone substitute material, promotes and supports migration, attachment, proliferation, and also phenotypic expression of bone-forming cells from the surrounding host tissue. It is directly dependent on the porosity and interconnectivity of the bone graft substitute, and also in a lesser extent by its chemical and physical properties of the substrate that promote adhesion and cell growth. By its definition osteoconductivity is a passive process. Osteogenesis is the process during which osteoprogenitor cells mature into osteoblasts, which subsequently mineralize and form bone tissue [8,53]. Besides biological properties, bone graft substitutes should offer optimal biomechanical strength, especially in those segments – such as femur and tibia – that are under high weight-bearing loads, or forearm and humerus that are subjected to high torsion forces. The biomechanical strength is a result of a complex interplay between the bone and the bone graft substitute material. In an ideal situation a bone substitute material may offer the same biomechanical strength as the bone being replaced. Bone reconstruction by materials with above mentioned properties is unfortunately typically limited to defects less than 6 cm in size [54].

The current golden standard treatment of critical-sized bone defect is autogenous bone grafting [3,55]. The application of autologous cancellous and corticocancellous graft, even vascularized, segmental bone grafts, when the defect exceeds some centimetres, allows in many cases the best management of the issue. However the complication rate in autogenous bone grafting is as high as 30% and may include donor site morbidity, pain, paresthesia, prolonged hospitalization and rehabilitation, increased risk of deep infection,

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11 hematoma, inflammation, cosmetic defects, tumour transplantation, and restricted availability [3,52]. Moreover, the morbidity of the procedure remains a drawback that has driven development of new strategies based on bone engineering, which may induce bone formation with less morbidity and similar efficiency [56]. Alternative strategies include the use of allografts or demineralized bone matrix or the use of synthetic material. Still those strategies rely on the possible migration of the surrounding cells into the implantation site to allow local bone regeneration.

Approximately one-third of the bone grafts used in North America are allografts [57]. It is an attractive alternative to autogenous bone as it avoids donor site morbidity, the allograft has osteoconductive properties and is available in various shapes and sizes. Allograft offers optimal osteoconductive and biomechanical characteristics due to its three-dimensional structure similar to that of human bone and the presence of collagen type I. Unfortunately, allografts carry the risk for viral disease transmission, bacterial infection, and immune rejection. Though the risk of transmission of disease is much lower than with blood products, it is still possible. Additionally, some existing tissue processing techniques which try to eliminate the risk of disease transmission can alter the graft’s biomechanical and biochemical properties [58].

Due to the expected shortage of allograft bone and the risk of virus transfer there has been an increased interest in bone substitutes. A large number of bone-graft alternatives are currently commercially available for orthopaedic use. One of them which is directly based on natural bone tissue is demineralised bone matrix (DBM), which is produced through decalcification of cortical bone. It contains collagen type I, non-collagenous proteins, and a small amount of osteoinductive growth factors such as bone morphogenetic proteins (BMPs) [57]. Since its antigenic surface is destroyed during demineralization, DBM does not evoke any appreciable local foreign-body immunogenic reactions. DBM is available on the market in form of granules, strips, putty, gel and freeze-dried powder. However, DBM has demonstrated a lower osteoinductive capacity compared to autologous bone grafts and has shown a high and questionable variability of the concentration of growth factors which depends on the manufacturer and manufacturing process [8].

Another approach is distraction osteogenesis (DO). It is a surgical technique, invented by Ilizarov, which features de novo bone formation between vascular bone surfaces created by osteotomy and gradual distraction [59]. For instance in alveolar ridge

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augmentation DO is accomplished when the coronal segment of the alveolar ridge is separated from underlying bone (while remaining attached to the lingual periosteum) with a saw or piezotome, and the two bone fragments are then slowly separated with a sliding bone plate [60]. It has been shown that distraction is able to induce new bone formation that is similar to natural bone. DO has been widely utilized to treat leg length discrepancy, deformity, non-union, osteomyelitis, and bone loss [59]. The main disadvantages are that the devices used in this technique for bone separation are cumbersome, technically difficult to apply, and require frequent manipulations during treatment. Additionally patients experience a lot of pain during treatment and are prone to bacterial infections [61].

Recently, autologous platelet-rich plasma (PRP) in the form of activated platelet gel and recombinant BMPs have been used with various results for healing wounds, ulcers, fractures and in maxillofacial setting. PRP is made by isolating a concentration of platelets from the patient's own blood. Platelet gels contain growth factors and function as osteoinductive agents. They can thus play a key role in bone formation and maturation of osseous fusions. Several studies demonstrated that better and stronger bone was yielded with the use of platelet gel combined with particulate bone allografts or other bone-filling materials as compared to reconstruction with conventional methods, and that bone density and growth were enhanced [62–65].

1.3.2 Overview of tissue engineering approach

More than half a million patients receive bone defect repairs each year in U.S. which costs more than $2.5 billion [66]. Due to increased life expectancy this figure is expected to be much higher in next decade. According to the Orthopaedic Industry Annual Report and GlobalData’s report, released in ORTHOWORLD, worldwide orthopaedic product sales exceeded $43.1 billion in 2012, increasing from $15 billion in 2002. Additionally, a recent report from Medtech Insight revealed that bone graft substitute products for spinal fusion totalled approximately $177.1 million in 2010 in Europe and are predicted to increase at a compound annual growth rate of 17.3% [67]. Currently, many countries worldwide are experiencing an exceedingly high demand for functional bone grafts. The use of bone grafts in the clinical practice is on the one side characterized by high success percentage but there are several major inconveniences. Complications or non-unions are common especially in large shaft reconstructions. Moreover, the substitution of the damaged site with host bone is frequently not fully achieved. This may lead to

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13 a secondary fractures, reported as high as 60% at 10 years. This is associated with a multitude of biologic processes such as osteoclastic resorption, increase in microfracture prevalence and decrease in bone mineral density within the allograft cortex, as well as the cancellous bone [6].

Tissue engineering comprises both in vivo and ex vivo strategies. The in vivo approach is to systemically or locally stimulate bone healing and regeneration. For instance the use of osteoinductive scaffolds is widely clinically applied. The classical ex vivo approach uses combinations of scaffolds with growth and differentiation factors and with skeletal and endothelial progenitor cells to prefabricate tailored bone constructs for transplantation. Unfortunately none of ex vivo procedures has yet reached clinical routine or proven to be superior to the autologous transplants [17]. The classic bone tissue engineering approach consists of several key elements: a biocompatible scaffold that closely mimics the natural bone extracellular matrix niche, osteogenic cells to lay down the bone tissue matrix, morphogenetic signals that help to direct the cells to the phenotypically desirable type, and sufficient vascularization to meet the growing tissue nutrient supply and clearance needs [66]. It involves interdisciplinary expertise in biomaterials, biomechanics, and cell-molecular biology to produce therapies that may either replace and/or regenerate bone. The implantation of exogenous pluripotent stem cells has been central to modern regenerative medicine approaches. Growth factors and scaffolds provide essential biological roles for the augmentation of stem cell-based therapeutics. They might also offer stand-alone therapeutic options [3,68].

1.3.3 Porous scaffolds for bone tissue

The key factors for an ideal scaffold for bone tissue engineering are: (i) pore size >100 μm; (ii) interconnected open porosity for in vivo tissue ingrowth; (iii) initial strength for safe handling during sterilizing, packaging, transportation to surgery, as well as survival through physical forces in vivo; and (iv) sterile environment for cell seeding; (v) sufficient mechanical strength for proper load transfer to the adjacent host tissue [69,70]. The extensiveness of a catalogue of requirements for bone substitute materials, however, demonstrates the complexity of the idea: scaffolds serve as temporary matrices for bone regeneration and their specific material properties are crucial for the successful outcome of the healing process [71]. The contemporary biomaterials for bone tissue regeneration can be classified broadly into inorganic and organic materials, which include both naturally

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derived and synthetic components [53]. The scaffolds for bone tissue engineering are made of metals, ceramic, polymers and composites. However, each of the material type possesses some distinctive problem that makes it less viable in comparison to bone autografts and allografts. In the case of metallic scaffolds it is toxicity caused by ion release, for polymeric scaffolds it is poor strength, for ceramics it is friability, for composites it is uncontrollable degradability, and so on [67]. The predominant group of scaffolds used in bone defects repair are ceramic scaffolds.

Calcium phosphate (CaP) synthetic substitutes are osteoconductive, but they are not osteoinductive unless growth factors, BMPs, or other osteoinductive substances are added to create a composite graft. Most calcium phosphate ceramics currently under investigation are synthetic composites of hydroxyapatite (HAp), tricalcium phosphate (TCP), or both. Because of the wide difference in resorption rates and porosity between TCP and HAp, a mixture of the two is clinically favourable [72]. Microwave sintered 3D printed β-TCP scaffolds with >60% porosity not only facilitated osteoblast activity but also aided in new bone formation in the pores [73]. Synthetic calcium phosphate ceramic based on hydroxyapatite (Ca5(PO4)3OH) is bioactive in the sense that it is a non-toxic compound

and interfacial bonds are able to develop between HAp and the living tissues leading to enhanced mechanical strength of the overall structure. Porous particles of HAp (average pore size 150 µm, porosity 70%) and porous coral-replicated HAp (exoskeletal microstructures of calcium carbonate of corals converted into hydroxyapatite by hydrothermal chemical exchange) blocks (average pore size 230 µm, porosity 66%) were used for delivery of BMP-2 in a rat ectopic model and induced direct osteogenesis (without preceding cartilage formation) [74]. While HAp has excellent bioactivity, its poor mechanical properties compared with bone have hindered its clinical application [75].

Recent studies have suggested that nanocrystalline HAp (nHAp) powders exhibit improved sinterability and enhanced densification due to their higher surface area, which may improve fracture toughness as well as other mechanical properties. To capitalize on its advantages and simultaneously overcome the drawbacks nHAp is combined with various types of polymers to generate composite materials that can be used for osteoconduction in the field of orthopaedic surgery [76]. Numerous studies showed the possible incorporation of nHAp into different polymeric materials, such as chitosan, collagen, PLGA, poly(L-lactic acid) (PLLA), hyaluronic acid, gelatin, poly(ε-caprolactone)

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15 (PCL), etc. The nanoscale features of HAp particles induce advantageous cellular responses when compared with micro-sized HAp particles [76–78].

Bioglass based bioresorbable scaffolds have been explored over the years since the development of resorbable 45S5 Bioglass®. The advantages of the glasses are an ease in controlling chemical composition and, thus, the rate of degradation which make them attractive as scaffold materials [79]. When tested in vitro, a 70% porous 3D bioglass scaffold with 300 to 400 μm pore size showed hydroxycarbonate apatite layer formation on its surface that significantly enhanced osteoblast activity [80]. In a rabbit tibia model, Goodridge et al. observed bone ingrowth into the pores of an apatite–mullite glass–ceramic scaffold prepared by selective laser sintering after implantation for 4 weeks [81].

Ceramic TiO2 has been shown to have excellent biocompatibility, especially when

in contact with bone tissue. This material has been implanted in vivo with few signs of inflammatory responses. TiO2 has shown to be biocompatible, enhance bone and vascular

ingrowth and to have a certain degree of bacteriostatic effect [82,83]. Several attempts have been made to manufacture implants and scaffolds based on its superior biocompatibility [84–86]. Mostly, TiO2 scaffolds were prepared by polymer sponge replication method [87].

However, one of the major obstacles by using TiO2 as a bone support is its limited

mechanical properties which result in having inadequate strength to serve as scaffolds in applications where higher loads are present [86]. Nonetheless, the animal trials on minipigs and rabbits showed that TiO2 scaffolds had excellent osteoconductive capacity and

provided a favourable environment for bone ingrowth [88,89].

Another approach is combining organic and inorganic materials which facilitates manufacturing of biocompatible scaffolds with appropriate mechanical properties for osseous defect sites. PCL microstrands have been combined with electrospun collagen nanofibers without compromising the cell adhesive properties of collagen and the mechanical strength of PCL [90]. The blending of chitosan and hydroxyapatite in scaffolds has resulted in materials with mechanical properties, porosity and bioactivity to support ingrowth of cells and new bone formation [91]. Other examples of recent composite biomaterial platforms include collagen and HAp [92], polyglycolide and βTCP [93], as well as a particularly novel combination of polyethylene glycol, PCL, collagen and nHAp [94].

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1.3.4 Therapeutic strategies in osteomyelitis treatment

Treatment of osteomyelitis infections is challenging due to a variety of factors, including the poor bioavailability of antibiotics in bone tissue, rising antibiotic resistance in bacterial pathogens, and the biofilm-like properties of the infection [30]. Acute osteomyelitis is often treated with systemic antibiotic administration. The standard treatment of the chronic osteomyelitis includes both adequate surgical debridement and systemic antibiotic therapy (with antibiotics like Vanc, ciprofloxacin, and Gent) [95]. A typical protocol of antibiotic treatment for chronic long-bone osteomyelitis in adults is intravenous administration of antibiotics for 4-6 weeks [96]. The time requirement is based on the fact that it takes that long for debrided bone to be protected prior to revascularization. Nonetheless the duration of therapy remains empiric. There are no clinical studies or documented records indicating the superiority of the 4-6 week course of antibiotics over other durations. Antibiotics are started empirically after cultures have been obtained, usually at the time of debridement in order to tailor accordingly the antibiotic regimen. What is essential to understand it is that antibiotic regimens are insufficient without adequate debridement, soft tissue coverage, and bone vascularity [97].

Unfortunately, the parenteral or oral application of the drugs increases chances of ototoxicity, nephrotoxicity and other untoward or adverse hypersensitivity reactions due to needed high concentration of antibiotic. This is due to elusive clinical response to antimicrobials, caused by numerous factors like poor penetrability of the drugs preventing the drug to attain required minimum inhibitory concentration level in bone tissue and prolonged course of therapy. Therefore, when selecting an antibiotic regimes such parameters must be taken into consideration like penetration into osseous tissue, effectiveness against the likely pathogens, frequency of dosing, drug toxicity and treatment cost. The available evidence indicates that for most antibiotics its concentration in serum corresponds with the one in bone [98]. Antibiotic treatment to prevent or eradicate a pathogenic infection is quite likely to induce concomitant damage to the commensal microbes of the human body, which can lead to both transient and persistent changes in host health and physiology, including dietary and nutritional processing, prevention of pathogen invasion, and immune system maturation [99].

Staphylococcus aureus strains are increasingly methicillin-resistant due to the

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17 (MRSA) and the recent emergence of community-associated MRSA, which have become a major cause of aggressive bone and joint infections in children and adults [98]. Beta-lactam antimicrobials remain the drugs of choice for nonallergic patients with methicillin-susceptible Staphylococcus aureus (MSSA) infections [28]. Although considered inferior to parenteral β-lactam therapy, there are several parenteral and oral options for treating MSSA osteomyelitis in patients allergic to penicillins and cephalosporins [98]. Glycopeptides like Vanc and drugs like daptinomycin and teicoplanin which form the inevitable part for treating MRSA or other anaerobic infection in the bone have poor bone penetrability [100]. Nonetheless, despite poor results in animal models, intravenous Vanc and teicoplanin have been used successfully for treating both MSSA and MRSA infections. Use of Vanc for treatment of osteomyelitis has increased dramatically with the emergence of MRSA, which now comprises the majority of S. aureus infections found in hospitals [98].

The goal of debridement is to reach healthy, viable tissue. Adequate debridement may leave a large bone defect known as dead space. Those spaces are filled with tissue flaps, free flaps, bone grafts or temporarily with poly(methyl methacrylate) (PMMA) beads loaded with gentamicin [97].

1.3.5 Localized drug delivery in osteomyelitis

Management of chronic osteomyelitis with the local delivery of the antimicrobial agent is novel therapeutic approach. It assures elevated antibiotic concentrations at the site of infection without systemic toxicity. Local treatment increases success rates and can be performed with different antimicrobial bone graft substitutes. Numerous studies have shown that direct application of antibiotic solutions to a wound does not help to clear infection [101]. On the other hand systems that allow gradual release of antibiotic over time are much more beneficial. Two main types of drug delivery systems can be distinguished: non-biodegradable and biodegradable.

Antibiotic-impregnated non-biodegradable beads mainly from PMMA have been widely used for the local administration of antibiotics and dead space management for more than 30 years. PMMA is available in two forms – bead cements and bead chains impregnated with antibiotic which are used for management of arthroplastics and musculoskeletal infections, respectively. The main advantage of PMMA is that it does not trigger any immune response from the host and the form of a bead confers a wide surface

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18

area, allowing rapid release of the antibiotic. The most commonly used antibiotics in bead have been the aminoglycosides, Gent or tobramycin [100]. A disadvantage of PMMA is that the material is non-biodegradable, making subsequent invasive procedures necessary to remove the implant in many cases [102].

In recent years, various biodegradable delivery systems have been developed and evaluated for local delivery of antibiotics in the treatment of bone infection. For instance calcium sulphate has been investigated in latest studies conducted by McConoughey. The performance of synthetic, high purity calcium sulphate as a carrier of antibiotics in orthopaedic infection has been equivalent, and potentially superior to PMMA [103]. Both the antibiotic impregnated materials have displayed similar bactericidal properties and similar kind of Vanc elution. Similar study confirmed the usefulness in chronic osteomyelitis treatment of tobramycin loaded biodegradable calcium sulphate carriers. The study showed that with supplementation of 10 to 20 gram of tobramycin impregnation, chance of renal toxicity is negligible in healthy persons. On the other hand, there was a serious concern for the patients with renal failure [104].

Collagen has been reported as a well-tolerated and completely biodegradable material. There have been several studies recently that show promising results in open fracture management by using antibiotic loaded collagen sponges locally along with parenteral therapy. Although the Gent-loaded collagen sponge provided high antibiotic concentration at the site of implantation, an additive effect of systemic antibiotic treatment was attained [105,106]. Sternal wound is a serious complication following cardiac surgery and Gent-loaded collagen sponge can be a good alternative for treatment of such wound [107,108].

Biodegradable synthetic polymers are known to be used in surgical operation for long. However, these materials are also gaining popularity as local antibiotic delivery system among the clinicians for their lasting effects and antibiotic elution with increased penetrability in bone and soft tissue infection [100]. The degree of biodegradability can vary from weeks to months, allowing variable types of infections to be treated. In addition, there may be decreased need for reconstruction of tissue defects as the host soft tissue can fill in while the beads are slowly absorbed. Biodegradable implants in comparison to PMMA cements show almost no retention of antibiotic [109]. Synthetic polymers are highly compatible with a number of antibiotics like ampicillin, Gent and polymixin-B.

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19 Polymers of lactide and glycolide are generally used for this purpose [100]. PLGA fully degrades within 1-6 months (depending on the composition), and its release profiles, whether primarily determined by diffusion or bulk erosion, vary [99]. Orhan at al. investigated PLGA microspheres with teicoplanin in tibial MRSA osteomyelitis. The results on rat animal model showed that the systems were clinically effective [110]. Among other materials as synthetic polymers are polyhydride and polycaprolactones. Copolymers of lactide and glycolide (90:10) are also used for their better compatibility with antibiotics like tobramycin, clindamycin and Vanc [111]. These materials degradevery slowly over a long period of time only at physiological pH and thus can provide sustained release of antibiotics. Not only that, drug elution kinetics from these synthetic polymers can be modulated by changing physical, biochemical and molecular structural properties of the polymers [112]. PLGA nanoparticles can also provide high cellular accumulation of antibiotics and effectively kill intracellular S. aureus [113].

Bio-absorbable gels containing antibiotics were found to offer faster recovery rate in comparison to standard bone cements. Recent studies show significant reduction of infection rate using bio-absorbable gels containing antibiotics like Gent and Vanc or with no antibiotics at all (P < 0.001) in an experimental open fracture model [114].

1.3.6 Therapeutic strategies in osteoporotic fractures treatment

Over the last decades, there have been substantial improvements in the treatment of osteoporosis. Nonetheless osteoporotic fractures are still a major clinical challenge in the elderly population due to impaired bone healing. Osteoporotic bone has reduced mass which leads to a smaller bone/implant contact area and consequently decreases the holding power for implants. Increased bone fragility predisposes to high rate of implant fixation failure and less than optimal environment for bone formation, which leads to prolonged healing time and increased risk of non-unions. Bone deposition by osteoblasts is in imbalance with osteoclast-mediated bone resorption, ultimately resulting in a net loss of bone over time. Additionally, age-related osteoporosis is mediated by decreased number and activity of osteoblastic cells, which leads to decline in bone formation and deterioration of bone microarchitecture [46,115]. Recent studies have recognized changes in the survival, activity and function of mesenchymal stem cells in aged and osteoporotic bone, which could be correlated to the diminished callus formation in osteoporotic fracture [116,117]. Levels of various systemic hormones and factors play important roles in regulating bone

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20

remodelling, including estrogen, PTH, Calc, glucocorticoids and vitamin D metabolites. Alterations in the stimulatory factors released locally from bone matrix during resorption or fracture, e.g. macrophage colony-stimulating factor, transforming growth factor (TGF)-β, platelet-derived growth factor, fibroblastic growth factors, BMPs, and insulin-like growth factor (IGF)-1, can affect both osteoblast and mesenchymal stem cells behaviour [115].

Osteoporosis therapies fall into two classes, antiresorptive drugs, which slow down bone resorption, and anabolic drugs, which stimulate bone formation [118]. Current FDA-approved pharmacologic options for osteoporosis treatment are bisphosphonates (alendronate, ibandronate, risedronate and zoledronic acid), Calc, estrogen agonist/antagonist (raloxifene), estrogen and hormone therapy, parathyroid hormone (teriparatide) and RANKL inhibitor (denosumab) [47].

Bone-anabolic drugs that build up new bone, rather than preventing its loss, are limited to the full-length parathyroid hormone (PTH 1-84) or its N-terminal fragment, teriparatide (PTH 34). Both are given subcutaneously, but transdermal forms of PTH 1-34 are in development[118].

Among drugs preventing bone resorption, bisphosphonates, with their high affinity for bone and long safety record, constitute the largest class. Bisphosphonates are known to increase bone mineral density by inhibiting osteoclast-mediated bone resorption and thereby reduce the risk of fractures. The nitrogen-containing bisphosphonates act by binding and blocking an enzyme in the 3-hydroxy-2-methylglutaryl-CoA reductase pathway (i.e. the mevalonate pathway), thus blocking the prenylation of small GTPases that are essential for osteoclast functionality and survival [115,119]. Unfortunately in a systemic oral administration of bisphosphonates there are several side effects, such as gastrointestinal problems such as erosions and ulcers in the stomach and small intestine, problems with renal function, atypical femur fractures, esophageal cancer, elevated serum creatinine levels [118,120]. There have been rare reports of osteonecrosis of the jaw with long-term use of bisphosphonates [121]. Bisphosphonates can be given orally or intravenously, and are most widely used because they can be inexpensive and used across a broad spectrum of osteoporosis types, including postmenopausal, male, and steroid-induced osteoporosis, as well as Paget’s disease [118]. One of the bisphosphonates, Aln, reduces the incidence of spine and hip fractures by about 50% over 3 years in patients with

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21 a prior vertebral fracture or in patients who have osteoporosis at the hip site [51]. It also increases bone mass. The main advantage of bisphosphonates are the lasting effects even after treatment discontinuation. All of non-bisphosphonate medications produce only temporary effects which wane upon discontinuation of treatment (denosumab, hormone therapy etc.) [47].

The effect of Calc now thought to be dependent on Calc binding to specific receptors on osteoclasts cause inhibition of bone resorption and decreasing the activity of those resorptive cells. Calc also binds to specific receptors on cells of the monocyte−macrophage series, proposing that it also acts to decline in bone resorption by retarding recruitment of osteoclasts. However, the reduction in bone turnover associated with Calc treatment is much smaller than that seen with other antiresorptive agents [122]. Calc also seems to have an analgesic effect on acute painful vertebral fractures [123]. Salmon Calc is approved by FDA for the treatment of osteoporosis in women, at least 5 years after menopause where no alternative treatments are available. Calc reduces vertebral fracture occurrence by about 30% [47]. A nasal spray is the most popular formulation of Calc among elderly patients, owing to its ease of administration [120].

1.3.7 Localized drug delivery in osteoporosis

The most important factors which predispose the local drug delivery systems for osteoporosis treatment is the improvement of the bioavailability of the drug and pharmacokinetics. Regarding fracture repair the main focus is on the carriers of the antiresorptive drugs [115,124]. Processes used for the preparation of these carriers include multiple techniques such as emulsion solvent process, spray drying, emulsion polymerization, emulsion solvent diffusion, rapid expansion of supercritical solutions, and ionic gelation. These various techniques led to the obtaining of a multitude of pharmaceutical forms ranging from polymeric MPs and nanoparticles to nanocrystals and liposomes [124]. A critical issue in local drug delivery for enhancing fracture repair in osteoporotic bone is dosage. In the case of osteoporotic fracture repair, it is not yet clear which carrier properties and release profile would be most suited and efficacious in terms of rapid increase in bone quality, bone-to-implant contact and mechanical stability [115].

One of the local delivery strategies is impregnating the injectable bone cements with anti-osteoporotic drug. Panzavolta et al. prepared and utilized biomimetic cements at different gelatin concentrations (10, 15 and 20 wt%) to introduce disodium Aln up to

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a concentration of 25 mM. In vitro results of this study indicated that Aln promotes osteoblast differentiation and inhibits osteoclastogenesis and osteoclast function [125]. Boanini et al. developed and assessed HAp nanocrystals. The in vitro assessment of HAp nanocrystals loaded with Aln at 7 wt% showed a significant decrease in the number of osteoclasts (by 30%) and a significant increase in osteoblastic activity, revealed by a twofold increase in the synthesis of alkaline phosphatase, osteocalcin and type I collagen [126,127].

Another strategy is coating of the implants or scaffolds (in the case of large bone defects) with the drugs. BMP-2 as a model bone anabolic molecule and the influence of its mode of loading into a metal implant surface was assessed in a minipig model. An improved bone–implant contact was found in the case of titanium implants containing BMP-2 absorbed on a biomimetic CaP coating compared to BMP-2 absorbed directly on the titanium implants [128].

Collagen type I/chondroitin sulphate in solution, simvastatin–chitosan complexes and zoledronic stearate in acetone solution have all been coated on titanium and reportedly improved osteoporotic bone–implant interface integration. PTH was also introduced onto a biomimetic CaP coating using co-precipitation [115]. Anti-resorptive bone drugs have been combined as coating on implant surfaces. One of the studies concerning CaP coating on titanium implant impregnated with zolendronate showed a significant improvement of bone quality and preserving bone volume around dental implants [129]. In another study titanium implants were coated with mesoporous TiO2 films on which Aln and raloxifene

were immobilized. In experimental study on rats, the ultrastructural interface analysis revealed enhanced apatite formation inside the raloxifene coating and increased bone density outside the Aln coating [130].

Particles as a drug carrier are an attractive option for a versatile delivery system, which can be used as a stand-alone system or combined within cements, hydrogel or even coated on implants. Drugs embedded in organic or inorganic particles is on one hand protected from the biological surroundings until its release, which on the other hand can be controlled by suitable selection of material and particle size. Particles can be of macro-, micro- and nanosize [124].

Kim et al. prepared bioabsorbable CaP microspheres that have been incorporated with Aln through an in situ loading process. The evaluation of the biological activity

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23 showed that Aln-loaded CaP microspheres directly blocked osteoclast formation by releasing Aln to monocytic precursor cells and effectively inhibiting their differentiation into osteoclasts [131]. Calc has been encapsulated in polyethyelene glycol and hydroxypropyl-beta-cyclodextrin MPs for administration by the pulmonary route which increased significantly Calc bioavailability [132].

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24

Chapter 2. OPTIMIZATION OF TITANIUM DIOXIDE SCAFFOLDS

2.1 Introduction

Ceramic materials can be processed into porous scaffolds whose morphology, mechanical parameters and chemical composition resemble those of mineral part in bone tissue. Bioactive materials like hydroxyapatite (HAp) or resorbable tricalcium phosphate (TCP) are often used for scaffold manufacturing [78,133]. Nowadays, researchers turned also their attention to titanium dioxide (TiO2). It is a biocompatible and non-resorbable

material that has the ability to induce direct bonding with bone tissue [84,85,88].

This chapter is focused on manufacturing titanium dioxide scaffolds using the polymer sponge replication method in one-step sintering (OSS) and two-step sintering (TSS) conditions. Different temperatures and times of both steps were applied in order to optimize manufacturing process. To this end microstructure by scanning electron microscopy (SEM) and by micro-computed tomography (µCT) as well as mechanical properties of the scaffolds were assessed. Additionally, a statistical analysis was performed in order to find a correlation between microstructure parameters and compressive strength of TiO2 scaffolds.

2.2 Materials and methods

2.2.1 Powder preparation

TiO2 powder was cleaned to remove phosphate ions that are typically present on

the particles’ surface (due to the manufacturing process). TiO2 powder (350 g) was soaked

in 400 mL 1 M NaOH. First NaOH solution was added to 1000 mL glass flask, then TiO2

powder was gradually added, mixed intensively for 1 min and left for sedimentation (around 30 min). Subsequently supernatant was carefully removed and deionized water was added up to 1000 mL mark. Again the content of the flask was vigorously shaken and set aside, followed by replacing the water after full sedimentation. Whole procedure was repeated 6 times until milky supernatant developed. The powder was later rinsed with 550 mL of 0.1 M HCl and again washed with deionized water (until pH between 3.4 and

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25 3.5 was achieved). Excess water was removed, TiO2 paste was spread on filter paper and

dried for 24 h at 37 o_{C. Then dried mass was broken into small pieces and sieved using}

vibratory sieve shaker (Analysette 3, Fritsch GmbH, Idar-Obersteinm Gernany) at amplitude 2.0-2.5. A column of sieves with decreasing pore size was used to separate particles below 100 µm and the whole procedure was repeated three times in order to achieve better efficiency. Powder was stored in sealed plastic bags for further use.

2.2.2 Materials

TiO2 powder (Kronos 1171, Kronos Titan GmbH, Leverkusen, Germany),

polyurethane foams (60 ppi, Bulbren S Eurofoam GmbH, Wiesbaden, Germany), sodium hydroxide (1 M NaOH; AnalR NORMAPUR, VWR International AS, Leuven, Belgium) and hydrochloric acid (0.1 M HCl and 1 M HCl; R.P. Normapur AR, VWR International As, Oslo, Norway) were used in the experiments [134]. Figure 2-1 shows a flowchart outlining scaffolds’ preparation procedure.

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2.2.3 Slurry preparation

The ceramic slurry was prepared by gradual addition of 65 g cleaned TiO2 powder

to 25 mL of UHQ-water. TiO2 powder was mixed with UHQ-water at 1000 rpm speed

(Digital Programmable Omni Macro ES Homogenizer, Omni International, United States of America). First, approximately one-half of 65 g of the powder was added and stirred for 10 min. Afterwards the rest of the powder was added thereto and mixed for 15 min. To avoid the problem of the particles aggregation and to control the viscosity, the pH of the slurry was kept below 1.5 throughout the stirring by adding small portions of 1 M HCl (total volume of 2 mL). Stirring was continued for 2.5 h at a speed of 5001 rpm and the temperature of the slurry was kept at 15 o_{C. Immediately after stirring was finished, pH and}

viscosity of the slurry were measured. If necessary, viscosity of the slurry was adjusted with HCl or NaOH solutions, to assure the value around 40 Pa∙s at the lowest shear stress rate.

2.2.4 Polymer template coating and pre-coarsening

TiO2 porous scaffolds were manufactured by the polymer sponge replication

method. Polyurethane sponges were cut to cylinders (12 mm in diameter, 12 mm or 6 mm in height depending on intended future test) by punching with a metal stamp. Cylinders were then washed in tap water, detergent (Deconex, Burer Chemie AG, Reichshof, Germany) and alcohol (Absolute, Arcus, Oslo, Norway). Finally, templates were dried at 37 o_{C for 24 h and stored in polyethylene bags.}

The polymer foams were dipped into the ceramic slurry, excess of the slurry was removed by squeezing templates between two polyurethane foam sheets on a self-made device (Figure 2-2). Samples were placed on a porous ceramic plate and dried at room temperature for at least 24 h. Removal of the polymer material was performed by burn-out procedure. The samples were heated at a rate of 0.5 o_{C/min up to 450}o_{C and kept at that}

temperature for 1 h. To initially consolidate the green bodies, samples were further heated to 1100 o_{C at 1}o_{C/min, followed by cooling down to room temperature at 5}o_{C/min. Then}

samples were submitted to OSS or TSS sintering. From each slurry one batch of the samples was sintered as a reference at 1500 o_{C for 20 h with heating rate of 3}o_{C/min and cooling}