BONE TISSUE ENGINEERING IN TWO PRECLINICAL OVINE … · Bone Tissue Engineering in two preclinical...

BONE TISSUE ENGINEERING IN TWO

PRECLINICAL OVINE ANIMAL MODELS

Jan Henkel

Dr. med., MD

Submitted in fulfilment of the requirements for the degree of

Doctor of Philosophy (PhD)

School of Chemistry, Physics and Mechanical Engineering

Faculty of Science and Engineering

Queensland University of Technology

2017

Bone Tissue Engineering in two preclinical ovine animal models i

Keywords

Bone tissue engineering, segmental bone defect, large volume bone defect, tibial

bone, large animal, sheep, ovine, preclinical animal model, bone morphogenetic

protein, platelet rich plasma, scaffold, polycaprolactone, tricalcium phosphate,

alginate, hydrogel, melt electrospinning, direct writing, fused deposition

modelling

ii Bone Tissue Engineering in two preclinical ovine animal models

Abstract

From an orthopaedic surgeon’s point of view, segmental tibial defects with a substantial

loss of bone volume are amongst the most challenging bone defects encountered in

clinical practice. Bone tissue engineering applications to treat such defects have been

extensively investigated in the laboratories over the last three decades, but translation

into clinical trials or even routine clinical practice has not taken place on a large scale

yet. The research presented in this PhD thesis focusses on the generation of preclinical

evidence of the regenerative potential of novel tissue engineering applications in ovine

large animal models to bridge the scale-up gap between small animal models and clinical

translation. A comprehensive and in depth literature review on the state of the art of bone

tissue engineering in the 21st century is given first. Afterwards, a newly developed

spatio-temporal hybrid delivery system for BMP-2 (composed of melt electrospun

tubular mPCL-CaP-Scaffolds combined with medical grade alginate) is investigated in

QUTs well-established and extensively characterized 3cm-segmental tibial defect ovine

animal model. It is shown that this novel tissue engineering application is capable of

fully regenerating such extensive bone defects yielding mature bone tissue and full

restoration of load bearing capability. A detailed analysis of the results including

biomechanical testing, microcomputed tomography and various histological /

immunohistochemical staining methods shows results that parallel the outcome of

previous small animal studies in rats. The second part focusses on the establishment and

characterization of a new, large-volume 6cm-segmental tibial defect model building on

the expertise from the current 3cm tibial defect ovine animal model. In a pilot study, the

capacity of a tissue engineering construct (which was well-investigated in the 3cm defect

model before) consisting of mPCL-TCP scaffolds combined with PRP and rhBMP7 to

regenerate these large volume tibial defects is then analysed. It is shown that the tissue

engineering application is not capable of consistently regenerating these even more

challenging segmental bone defects. However, bone healing in this novel animal model

in the presence of a mPCL-TCP scaffold and with reduced doses of BMP-7 are

extensively analysed radiologically, histologically and immunohistochemically; further

characterizing this newly established animal model giving valuable insights for future

studies. All finding in this PhD thesis are presented in detail in the context of currently

available literature and implications for further studies and potential clinical translation

are discussed.

Bone Tissue Engineering in two preclinical ovine animal models iii

Table of Contents

Keywords .................................................................................................................................. i

Abstract .................................................................................................................................... ii

Table of Contents .................................................................................................................... iii

List of Figures ...........................................................................................................................v

List of Tables ......................................................................................................................... vii

List of Abbreviations ............................................................................................................ viii

Statement of Original Authorship ........................................................................................... ix

Acknowledgements ...................................................................................................................x

Chapter 1: Introduction ...................................................................................... 1

1.1 (Clinical) Background ....................................................................................................1

1.2 Bone Tissue Engineering ................................................................................................6

1.3 Thesis Outline .................................................................................................................7

1.4 Hypotheses ......................................................................................................................9

Chapter 2: Bone Regeneration based on Tissue Engineering Conceptions – A

21st Century Perspective ................................................................. 11

2.1 Bone biology .................................................................................................................13

2.2 Bone grafting and bone substitutes in the last 4000 years ............................................16

2.3 Bone substitute materials (BSM) ..................................................................................23

2.4 Three-dimensional scaffolds in bone tissue engineering ..............................................25

2.5 Additive manufacturing and Computer Aided Design – Game changers in the

fabrication of three-dimensional scaffolds ...................................................................35

2.6 Translating bone tissue engineering strategies from bench to bedside .........................39

Chapter 3: A novel spatiotemporal delivery system of rhBMP-2 to regenerate

critical sized tibial defects in an ovine large animal model .......... 57

3.1 Introduction ..................................................................................................................57

3.2 Materials and Methods .................................................................................................63

3.3 Results ..........................................................................................................................70

3.4 Discussion and conclusion ............................................................................................89

3.5 Acknowledgements.......................................................................................................97

iv Bone Tissue Engineering in two preclinical ovine animal models

Chapter 4: Establishment of a preclinical ovine animal model for the

treatment of large volume 6cm-tibial segmental defects .............. 99

4.1 Introduction .................................................................................................................. 99

4.2 Material and methods ................................................................................................. 103

4.3 Results ........................................................................................................................ 111

4.4 Discussion .................................................................................................................. 122

4.5 Acknowledgements .................................................................................................... 130

Chapter 5: Final Discussion ............................................................................ 131

Bibliography ........................................................................................................... 135

Appendices .............................................................................................................. 157

Appendix A Paper 1 ............................................................................................................. 157

Appendix B Paper 2 ............................................................................................................. 159

Appendix C Paper 3 ............................................................................................................. 161

Appendix D Paper 4 ............................................................................................................. 162

Appendix E Paper 5 .............................................................................................................. 164

Appendix F Paper 6 .............................................................................................................. 166

Bone Tissue Engineering in two preclinical ovine animal models v

List of Figures

Figure 1: Example of a severe high-grade open tibial fracture with

considerable, full-thickness contusion, abrasion, extensive open

degloving and skin loss (Type IO 4 according to AO soft-tissue

classification). ............................................................................................... 2

Figure 2: Hierarchical structural organization of bone. ...................................... 14

Figure 3: Clinical case combining the Masquelet-technique and the RIA-

system to treat a tibial non-union ............................................................. 22

Figure 4: Schematic illustrating the interdependence of molecular weight

loss and mass loss of a slow-degrading composite scaffold plotted

against time, which corresponds with tissue regeneration ..................... 28

Figure 5: For tissue engineering, the Valley of Death is the gap and

associated funding difficulties of taking tissue engineering

technologies to tissue-engineered products. ............................................. 40

Figure 6: Bone tissue engineering strategies .......................................................... 42

Figure 7: Load-bearing critical-sized ovine tibial defect model using

mPCL-TCP scaffolds manufactured by FDM. ....................................... 47

Figure 8: The use of mPCL-CaP scaffolds for spinal fusion. ............................... 50

Figure 9: Clinical case showing the craniofacial scaffold applications for

orbital floor fractures ................................................................................ 51

Figure 10: Clinical case of a 52 year old man with a malignant bone

tumour above his left hip ........................................................................... 54

Figure 11: The vascularised fibula transfer combined with bone tissue

engineering applications. ........................................................................... 55

Figure 12: Rat femoral defect model using hybrid delivery system of

rhBMP-2. .................................................................................................... 62

Figure 13: Representative radiographs at 4 and 12 weeks for rat femoral

defect model. ............................................................................................... 62

Figure 14: Representative images of tubular microfiber mPCL-scaffolds

surface-coated with CaP used in the study. ............................................. 64

Figure 15: Surgical procedure ................................................................................ 65

Figure 16: Representative clinical radiographic images at 3 and 6 months

after surgery. .............................................................................................. 71

Figure 17: Results of biomechanical testing at 6 months after surgery. ............. 72

Figure 18: Three-dimensional reconstructions of microcomputed

tomography (µCT)-scans ........................................................................... 73

Figure 19: Total Bone Volumes (TBV) (=bone volume over the complete

defect size) ................................................................................................... 74

vi Bone Tissue Engineering in two preclinical ovine animal models

Figure 20: Overview of results from histological stains and

immunohistochemical analysis of group III

(Scaffold+alginate+rhBMP-2-group, representative sample

specimen) ..................................................................................................... 75


immunohistochemical analysis of group II (Scaffold+alginate-

group, representative sample specimen) .................................................. 76


immunohistochemical analysis of group I (Scaffold only-group,

representative sample specimen) .............................................................. 77

Figure 23: Details of Haematoxylin-Eosin-Stain of representative samples

of group III (scaffold + alginate + rhBMP-2) .......................................... 81

Figure 24: Representative sample (group III) of anti-Collagen 1-antibody

IHC .............................................................................................................. 83

Figure 25: Representative samples of IHC using antibody against ALP ............ 84

Figure 26: Representative images (group III) for IHC using antibodies

against VEGF, CD31 and vWF ................................................................. 85

Figure 27: Representative images of IHC with anti-CD68 antibody ................... 86

Figure 28: Representative images of Tartrate-resistant acid phosphatase

(TRAP)-staining. ........................................................................................ 87

Figure 29: Representative images from IHC using antibody against BMP-

2&4 .............................................................................................................. 88

Figure 30: Direct comparison of results from rat femoral defect model (left

column) and ovine animal model (right column) .................................... 91

Figure 31: Surgical Technique for 6cm tibial defect animal model .................. 106

Figure 32: Conventional X-ray images and 3D-reconstructions of

mineralized tissue from µCT in 3-months-group (group I). ................ 113

Figure 33: Conventional X-ray images and 3D-reconstructions of

mineralized tissue from mCT in 12 month-group (group II ................ 114

Figure 34: Microcomputed tomography-results at three months post-

surgery ....................................................................................................... 116

Figure 35: Statistical analysis of microcomputed tomography at 12 months

post-surgery (A) and comparison between group I and II (B) ............ 117

Figure 36: Statistical analysis of biomechanical testing ..................................... 118


immunohistochemical analysis of group I (3 months time-point,

representative sample specimen) ............................................................ 120


immunohistochemical analysis of group II (12 months time-point,

representative sample specimen) ............................................................ 121

Bone Tissue Engineering in two preclinical ovine animal models vii

List of Tables

Table 1: Mechanical properties of compact (cortical) and spongy

(cancellous) bone ........................................................................................ 16

Table 2: Young’s modulus (GPa) (according to various levels of

architecture). .............................................................................................. 16

Table 3: Bone grafts and graft substitutes currently used in clinical

orthopaedic applications ........................................................................... 25

Table 4: Mechanical properties and degradation kinetics in relation for

porosity of composite scaffolds. ................................................................ 34

Table 5: Advantages of scaffolds designed and fabricated via additive

manufacturing ............................................................................................ 38

viii Bone Tissue Engineering in two preclinical ovine animal models

List of Abbreviations

2D Two-dimensional

3D Three-dimensional

ABG Autologous bone graft

AICBG Autologous Iliac Crest Bone Graft

AO Arbeitsgemeinschaft Osteosynthese

BMP Bone Morphogenetic Protein

BTE Bone Tissue Engineering

BV Bone Volume

CAD Computer Aided Design

CaP Calcium Phosphate

DCP Dynamic Compression Plate

ECM Extracellular Matrix

FDA Food and Drug Administration

FDM Fused deposition modelling

GBP Great Britain Pound

mPCL Medical grade Polycaprolactone

MSC Mesenchymal Stem Cell

PMMA Poly-Methyl-Methacrylate

PRP Platelet Rich Plasma

RIA Reamer Irrigator Aspirator

RM Regenerative Medicine

TBV Total Bone Volume

TCP Tricalciumphosphate

TE Tissue Engineering

TEC Tissue Engineering Construct

TE&RM Tissue Engineering and Regenerative Medicine

TM Torsional Moment

TS Torsional Stiffness

USD US Dollar

Bone Tissue Engineering in two preclinical ovine animal models ix

Statement of Original Authorship

“The work contained in this thesis has not been previously submitted to meet

requirements for an award at this or any other higher education institution. To the

best of my knowledge and belief, the thesis contains no material previously published

or written by another person except where due reference is made.”

Signature: QUT Verified Signature

Date: 2nd

February 2017

x Bone Tissue Engineering in two preclinical ovine animal models

Acknowledgements

I would like to gratefully acknowledge my supervisor Prof. Dietmar W.

Hutmacher for his great support (on a professional as well as personal level) and

excellent supervision of my work, for his continuous friendship and his inspiring

enthusiasm for the field of regenerative medicine. I had a great time at his institute

that I will always cherish. I would also like to thank Prof. Michael Schuetz for his

supervision, his constant support and feedback regarding my work.

Furthermore, I would like to thank Associate Prof. Mia Woodruff, Dr. Siamak

Saifzadeh and Dr. Roland Steck for their great help with all my research projects,

their expertise as well as personal friendships.

Thank you to Dr. Stephanie Fountain and Dr. Devakar Epari for analysing the

mechanical properties of the scaffolds and well as the entire construct in vitro of the

6cm tibial defect model. And for help with biomechanical questions in general.

A big thank you to all the members of the QUT Medical Engineering Research

Facility (MERF) for their great animal work with the sheep. Claudia, Andrew, Mark,

Anton: Your professional planning for my projects, assistance with the animal

surgeries, animal handling and postoperative care was remarkable. Ian, your

friendship and help in difficult times is highly appreciated.

Thank you to all the members of the Regenerative Medicine-Group and my

fellow researchers at the Institute of Health and Biomedical Innovation for working

with me on the projects, for scientific discussions as well as the new friendships. I

would especially like to thank Dr. Boris Holzapfel, Jeremy Baldwin, Dr. Arne

Berner, Dr. Cameron Black, Mohit Chhaya and Onur Bas.

Special thanks also to Joanne Richardson for helping with all administrative

work and being incredibly helpful with solving all of the small daily problems/issues

encountered during my PhD studies at QUT.

Thank you to the entire Biofabrication and Tissue Morphology group

(especially Flavia Medeiros Savi, Felicity Lawrence and Keith Blackwood) for

helping with the extensive histological and immunohistochemical analyses. You

have been great!

Bone Tissue Engineering in two preclinical ovine animal models xi

Thank you to my parents Ingelies and Wolfgang as well as my siblings Inga

and Nico for their encouragement, advice and support over the years.

I would like to thank my wife Hanna and daughter Edda for sharing the

wonderful experience of living and working in Brisbane with me. Thank you so

much for all your support, your understanding and all the love and happiness you are

giving me!

Bone Tissue Engineering in two preclinical ovine animal models xii

For Edda and Hanna

- My wonderful daughter and my beloved wife….my heroes -

Chapter 1: Introduction 1

Chapter 1: Introduction

1.1 (CLINICAL) BACKGROUND

The impact of musculoskeletal disorders on the individual health as well as the

socioeconomic situation is significant. Accounting for almost 25% of the total cost of

illness and up to 15% of the costs of primary care, they are the second most common

reason for consulting a general practitioner [1]. Numbers are predicted to grow with

increasing life expectancy, increasing incidence of lifestyle-related obesity, reduced

physical fitness and increased numbers of road traffic accidents [1].

Among musculoskeletal disorders, skeletal trauma and resulting fractures of

long bones are highly prevalent. It has been estimated that in the USA long-bone

fractures account for approx. 10% of all non-fatal injuries [2] and are the number one

category of injuries regarding inpatient expenditures [3]. In addition to direct medical

costs the costs for lost productivity due to workplace absences and (short-term)

disability represent a significant component of the burden of long bone fractures [4].

The tibia (shin bone) is the most commonly fractured long bone in the human

body with an annual incidence of 2 tibial shaft fractures per 1000 individuals [5]. The

average age of patients with tibial shaft fractures is approx. 40 years, with teenage

males being reported to have the highest incidence [6, 7]. In contrast to other

common fractures such as proximal femur (thigh bone) fractures, proximal humerus

fractures, distal radius fractures or pelvic fractures, tibial shaft fractures are not

regarded as predominantly osteoporotic fracture types and their prevalence does not

increase with age [6]. Tibial (shaft) fractures therefore have to be regarded as a

fracture type highly prevalent in a relatively young patient population with usually

good bone quality and relatively low numbers of comorbidities (diabetes,

osteoporosis, vascular diseases and so forth).

Nevertheless, the treatment of (open) tibial fractures still represents a major

clinical challenge and poses a significant risk of associated complications such as

infection and non-union [8-10]. Due to the thin anteromedial soft-tissue coverage of

the lower leg open tibial fractures (that is fractures with damage to/disruption of the

soft tissue above the bone and potential exposure of bone parts) are the most

2 Chapter 1: Introduction

common open fracture [11]. They are often associated with significant loss of bone

substance and severe damage to the surrounding soft tissue (Figure 1). High-grade

open tibial fractures (Gustilo-Anderson Type IIIb or IIIc [12]) carry an infection risk

of up to 25-50% [13, 14]

Figure 1: Example of a severe high-grade open tibial fracture with considerable, full-thickness

contusion, abrasion, extensive open degloving and skin loss (Type IO 4 according to AO soft-

tissue classification). (a-b) Schematic depicting extensive bony and soft tissue damage, (c) clinical

image of injury site, (d) Conventional plane X-ray image (lateral view) showing extensive bone

damage. Reproduced from Rüedi, Buckley, Moran – AO principles of Fracture Management, 2nd

edition, 2007, Thieme, Stuttgart, Germany. © Georg Thieme Verlag, all rights reserved.

The average time to union for uncomplicated tibial (shaft) fractures is approx.

one year, but complex cases can be much longer and require multiple surgical

interventions [15]. Tibia and femur have been reported to be the most common

fracture sites for development of pseudarthroses. Delayed union of bone or

development of pseudarthrosis (non-union of bone) is found in averagely 13% of all

tibial fractures [16]. However, studies have reported much higher non-union rates of

up to 50-80% depending on the injury type, presence of infection and surgical

treatment [10]. For example. high-grade fractures (AO Classification Type C2-3) or

associated open skin injuries >5cm were found to have significantly higher risks of

pseudarthroses (Odds Ratio 6.3 and 13.9, respectively) [16].


Sustaining a tibial fracture is in itself a significant and impactful event for each

patient individually as well as for the healthcare system in general. More than 70,000

hospitalisations, 800,000 office visits and 500,000 hospital days have been attributed

to closed tibial shaft fractures in the US annually [17]. However, the consequences of

suffering a severe (open) tibial fracture with threatening limb loss, potential

consecutive delayed bone healing or development of pseudarthrosis can be

devastating for patients, their families/social environment and the entire society (loss

of productivity, health care cost etc.). This is drastically illustrated by the fact that

only 28% of patients suffering severe open tibial fractures resume full function and

are able return to their previous employment [1]. Non-unions of tibial shaft fractures

are associated with substantial healthcare resource use, common and prolonged use

of strong opioids, and high per-patient costs [18]. Multiple surgeries are often

required, one study found that infected tibial non-unions required an average of 8.8

operations till healing vs. an average of 5 operations for aseptic tibial non-unions

[19]. The average total cost of treatment for each tibial shaft fracture that develops a

non-union has been estimated to be as high as GBP 21,183.05 compared to GBP

3,111 treatment costs in an uncomplicated clinical course of a tibial shaft fracture till

clinical and radiological union [20].

Reviewing the literature varying numbers of total treatment costs can be found

(due to differences in study designs, varying years of the studies, different currencies,

variations in treatment techniques and treatment costs and so on) (see for example [4,

18, 20, 21]). However, findings of all studies clearly point towards increased rates of

surgical interventions, higher total treatment costs, significant loss of productivity

and a significant decrease in quality of life for patient suffering from tibia fractures

with consecutive non-union. It can be concluded that tibial fractures and their non-

unions represent a significant burden for the individual patient as well as for the

healthcare system and society in general.

Fracture healing is a highly complex process involving multiple interdependent

cascades of events and therefore individual causes of delayed fracture healing or

progression to non-union can be challenging to identify and are often multifactorial

as well [22]. Fracture-specific risk factors include, amongst others, severe high-grade

fracture types, presence of large open wounds and extensive soft tissue damage as


well as the presence or development of infection [16, 23]. Due to an evident lack of

randomized controlled studies currently available literature is inconclusive with

regards to treatment-specific risk factors. It is uncertain which type of primary

surgical treatment (e.g. intramedullary reamed or non-reamed nail fixation, open

reduction and internal fixation with various plate and screw types, external circular

fixation) for tibial (shaft) fractures results in fewest re-operations, earliest bone union

and littlest rate of non-union [9, 24, 25]. Patient-specific risk factors such as Diabetes

mellitus, Anaemia, Hypothyroidism, Peripheral Vascular Disease, medication with

Steroids or Non-steroidal Anti-Inflammatory Drugs (NSAD) or Statins as well as

smoking (most well-documented modifiable patient-specific risk factor!) have been

linked to inhibition of fracture healing and potential progression non-union [23, 26].

A lack of a standardized and (clinically as well as scientifically) commonly

accepted definition for the term “non-union” is further complicating matters [27-29].

Compromising the comparability between different studies this lack of a clear

definition (amongst other factors) negatively affects evidence levels of available

literature. For this thesis, the author will follow the definition of the US FDA

defining non-union as incomplete fracture healing after 9 months following injury,

along with absence of progressive signs of healing on following serial radiographs

over the course of three consecutive months [30].

A number of different treatment options for (tibial) fracture non-unions can be

found in the literature [23, 31]. Adjunct therapies such as the application of Low

Intensity Pulsed Ultrasound (LIPUS) [32-34] or Teriparatide (recombinant human

parathyroid hormone) [23] may have beneficial effects on bone repair and fracture

healing, but few or no randomized control trials providing high level evidence on this

exist so far. However, tibial fracture non-unions remain a domain of surgical therapy.

Ruling out of or treatment of potentially present infection, local debridement,

adequate fracture stabilisation and bone grafting (when necessary) have been

proposed using both single or multi-staged procedures. The use of plate and screw

fixation, intramedullary nail fixation, nail dynamization or exchange nailing,

distraction osteogenesis via external fixation (Ilizarov method), induced membrane

techniques (Masquelet Technique) as well bone grafting and the use of

orthobiologics such as bone marrow aspirates, platelet-rich plasma (PRP) or growth

factors (e.g. Bone Morphogenetic Proteins, BMPs) has been reported [23, 31, 35-43].


Tibial fracture non-unions and consecutive (multiple) surgical interventions

often lead to substantial bone volume loss in the tibial shaft region. Along with

extensive loss of bone substance due to tumour resection (primary bone tumours

such as osteosarcoma [44] or secondary bone tumours) or revision surgery after

failed arthroplasties, these large segmental (tibial) bone defects are still a major

clinical challenge and frequently require the application of bone grafts and/or bone

substitute materials. The rangeof bone graft materials includes autologous bone

(from the same patient), allogeneic bone (from a donor), demineralised bone matrices

as well as a wide range of synthetic bone substitute biomaterials such as metals,

ceramics, polymers, and composite materials [45].

A total of approx. 3.5million bone grafting procedures are performed each year

worldwide with the market being estimated to be in excess of USD 2.5 billion with a

predicted increase of 7-8% per year [46]. Autologous bone grafts (ABGs), mostly

harvested from the iliac crest of the patient (Iliac Crest Autologous Bone Graft,

ICABG) or via Reamer-Irrigator-Aspirator-Systems (RIA-ABG) e.g. from the femur,

still represent the clinical gold standard bone graft [47-49]. However, graft volumes

are limited, an additional surgical procedure is required to harvest the ABG and there

is significant risk for donor site morbidity such as chronic pain or dysesthesia at the

donor site [50, 51]. Large volume bone defects (>5 cm) are most commonly treated

with vascularised fibula autograft[52] and the Ilizarov method [53-55] because of the

risk of graft resorption despite good soft tissue coverage [56]. Complications are

common for these procedures and the process can be laborious and painful for the

patient requiring external fixation for up to 1.5 years [57-59]. Alternatives to

autograft bone currently include allogenic grafts, xenografts or other bone substitute

materials/orthobiologics such as bone marrow aspirates, PRP, ceramics, polymers

and composite materials [45, 60]. For an extensive review on bone grafting

procedures and bone substitute materials the reader is kindly referred to Chapter 2.

Given the limitations of current available bone grafting procedures and the

increasing demand for bone repair in limb salvage surgeries (fracture non-unions,

bone tumours, revision surgeries of failed arthroplasties) TE and its application in

orthopaedics has received considerable scientific, economic and clinical attention

over the last three decades [45, 61, 62].


1.2 BONE TISSUE ENGINEERING

With the introduction of Tissue Engineering (TE) in 1988 and its clinical

brother Regenerative Medicine (RM), hopes were high that we would soon be

“pulling engineered organs out of the petri dish” [63]. Although we are not (yet) able

to engineer entire organs in a petri dish, innovative and exciting new bone tissue

engineering applications are trialled in laboratories and preclinical animal studies,

some of which have already been used in humans [45, 62, 64]. Further details and a

comprehensive literature review on Bone Tissue Engineering and its applications can

be found in Chapter 2.

Despite increasing research expenditures yielding numerous discoveries and

innovations in the field of bone tissue engineering, a large scale translation of these

novel techniques from bench to bedside has still not taken place. There is a stark

contrast between the amount of tissue engineering research expenditures and the

resulting numbers of products for clinical application over the last decades [65].

Many new ventures “die” in the “Valley of Death” between scientific technology

development and actual commercialization of the application due to technical

challenges, business challenges and/or philosophical challenges [65-67]. As recently

discussed, one of the critical aspects to overcome current challenges is the need for

early trans-disciplinary communication and collaboration in the development and

execution of research approaches [68]. The paucity of engagement with the clinical

community has been identified as a key contributor to the lack of commercially

successful products [69, 70]. The need to address the shortage of sustained funding

programs for multidisciplinary teams conducting translational research was regarded

equally important [68].

Hollister has pointed out that “defining specific clinical target applications [for

tissue engineering approaches] remains likely one of the most underestimated

challenges in translating tissue engineering research into tissue-engineered products

“[65]. It is imperative to assess the clinical demands to achieve a broad and

optimised range of clinical applications for the specific tissue engineering approach

to be translated. Reviewing Chapter 1.1 and taking the above discussed implications

of tibial fractures/tibial non-unions into account, it can easily be concluded that tibial

segmental defects still represent a major clinical challenge in orthopaedics with clear

need for improvement of treatment options.


Having assessed the current clinical situation and clinical demand for the

application of novel bone tissue engineering strategies in the treatment of segmental

tibial defects, the “Centre of Regenerative Medicine” (which includes members from

the fields of engineering, cell biology, chemistry, clinical and veterinary medicine)

has defined tibial defects as a target for clinical translation. Over the last decade the

Hutmacher group at QUT has developed and pursued a rationale and road map of

how a multidisciplinary research team can address the current challenges and

successfully translate orthopaedic bone engineering applications from bench to

bedside [71]. Next to the actual development of novel TE applications based on

sound scientific studies, one of the greatest difficulties in bridging the Valley of

Death is to develop good manufacturing processes, scalable designs and to apply

these in preclinical studies. Not only has the group investigated numerous high

impact TE applications in the laboratory, they have also developed a highly-

standardized and fully-characterised ovine large animal model for preclinical testing

of the regenerative potential of such applications in critical-sized segmental tibial

defects [71-78]. This model has not only generated a series of highly cited

publications, but also has attracted large interest in the medical device industry to be

used as a preclinical test bed for their bone graft products under development. The

model enables control of experimental conditions to allow for direct comparison of

products against a library of benchmarks and gold standards that have been

developed over the last 10 years.

1.3 THESIS OUTLINE

The preclinical tibial defect model developed at QUT is one of the few

available models internationally, which is suitable from both reproducibility and cost

point of view for the evaluation of large segmental bone defect repair technologies in

statistically powered study designs. The research is focused on generating preclinical

evidence for the efficacy and safety of novel tissue engineering applications that will

underpin the future clinical evaluation of such technologies and ultimately their

potential translation into routine clinical practice.

Chapter II of the thesis provides a state of the art review on bone tissue

engineering from a biomaterial science, tissue engineering and regenerative medicine


(TE&RM) as well as a clinical point of view. Furthermore, an overview over past

and current studies of the Centre of Regenerative Medicine is given.

Over the last decade a 3cm critical-sized defect model in sheep tibiae was

established and biomechanically, histologically and immunohistochemically

characterized. This animal model is used to evaluate different biomaterials based on

TE&RM-rooted bone tissue engineering concepts: So far, the regenerative potential

of different types of mPCL-TCP scaffolds [76], of mPCL-TCP scaffolds in

combination with Mesenchymal Stem Cells (MSCs) or different dosages of rhBMP-7

(3.5mg and 1.75mg, respectively) [75, 78], combined with autologous vs. allogenic

mesenchymal progenitor cells [74], combined with osteoblasts from the axial

skeleton vs. osteoblasts from the orofacial skeleton [72] or combined with a delayed

injection of allogenic bone marrow stromal cell sheets [73] has been analysed.

Chapter III analyses the regenerative potential of a novel spatio-temporal delivery

system for rhMBP-2 in the critical sized 3cm tibial defect ovine animal model.

The 3cm critical-sized tibial defect model is now well established at QUT and

many different tissue engineering approaches have already been analysed using this

test bed. However, bone substance defects encountered in clinical practice are often

of larger volumes in nature, especially after multiple surgical interventions for non-

unions, after tumour removal or in the context of revision surgeries for failed

arthroplasties. In order to reflect the clinical situation even better, Chapter IV reports

on the establishment and characterisation of a larger volume, 6cm tibial defect ovine

animal model to further investigate the bone regeneration potential of various TE

approaches under well characterised and highly standardised conditions. In order to

minimise potential confounding variables in this pilot study, a study design similar to

the 3cm tibial defect was chosen mPCL-TCP scaffolds loaded with PRP and rhBMP-

7.


1.4 HYPOTHESES

Hypothesis I:

It was hypothesized that the application of a spatio-temporal delivery system

composed of a medical grade PCL-scaffold designed and fabricated via melt

electrospun writing and CaP-coating combined with medical grade alginate and

recombinant human bone morphogenetic protein 2 (rhBMP-2) developed in a small

animal model would be transferable into a preclinical ovine large animal model.

Furthermore, it was hypothesized that the application of such hybrid delivery system

in a large animal model would lead to bone regeneration equal to the previous results

from the rat femoral defect model.

Hypothesis II

It was hypothesized that it would be possible to establish a larger volume tibial

segmental defect model (6cm length) building on the expertise from the current well-

established ovine 3cm tibial defect model. Furthermore, it was hypothesized that the

application of the combination of mPCL-TCP scaffolds with PRP and rhBMP7

(which was well-investigated in the 3cm defect model before) would also have the

regenerative potential to bridge these large volume tibial defects.

Chapter 2: Bone Regeneration based on Tissue Engineering Conceptions – A 21st Century Perspective 11

Chapter 2: Bone Regeneration based on

Tissue Engineering Conceptions

– A 21st Century Perspective

In 2007 Chris Mason proposed two distinctly different phases in Regenerative

Medicine/Tissue Engineering in analogy to the changes the world wide web had

undergone before [79]: A research intensive phase of regenerative medicine 1.0

(RegenMed 1.0) from 1985-2002 which was all about fundamental research and

scientific discovery, and had little focus on translation into applicable products. He

heralded the era of Regenerative Medicine 2.0 (RegenMed 2.0) from 2006 onwards,

with the focus almost exclusively on the translation of research into commercially

successful products and an emphasis on the use of human embryonic stem cells

(hESCs) for future regenerative medicine applications. Since then, the potential of

ESCs for the use in regenerative medicine has been discussed extensively and Mason

& Dunhill have comprehensively reviewed the value of autologous and allogeneic

cells for regenerative medicine in 2009 [80]. However, when looking at clinical

translation (especially for hard tissues such as bone) cell-based therapies have so far

largely failed from both a clinical and economical point of view [81, 82].

Additionally, the pragmatic approach of RegenMed 2.0 to focus on clinical

translation and large scale commercialisation does not allow incorporating

Personalised Medicine approaches in order to focus on the distinctly different

prerequisites in each individual patient in need of tissue engineering strategies.

After 17 years of RegenMed 1.0 and another 7 years of RegenMed 2.0 versions we

herein propose that the era of RegenMed 3.0 has begun. The phase of RegenMed 2.0

was mainly focused on the translation of scientific discoveries into routine clinical

practice from the stem cell biology point of view with large scale commercialisation

in mind. RegenMed 3.0 takes a more holistic approach to tissue regeneration

combining experiences in stem cell biology with the identification of specific clinical

target applications taking into account clinical demand and practicability of the

techniques as well as regulatory and economic factors and patient specific

requirements. The era of RegenMed 3.0 will encompass a significant step forward in

terms of personalised medicine. The complexity and great variety of large bone

12 Chapter 2: Bone Regeneration based on Tissue Engineering Conceptions – A 21st Century Perspective

defects require an individualized, patient-specific approach with regards to surgical

reconstruction in general and implant/tissue engineering selection in specific. We

advocate that in RegenMed 3.0 bone tissue engineering and bioengineering

technology platforms, such as additive manufacturing approaches, will be utilised

substantially in bone grafting procedures to advance clinical approaches in general

and for the benefit of individual patient in particular.

This review will describe the state of the art of the bone tissue engineering field and

present a perspective of its role in Tissue Engineering & Regenerative Medicine 3.0.

Reviewing the field it can be summarized that over the last ten years remarkable

progress has been made in the development of surgical techniques for bone

reconstruction. Although these sophisticated techniques have transformed

reconstructive surgery and significantly improved clinical outcomes, they have

already reached a number of their practical limits to further improve healthcare

outcomes. Today major reconstructive surgeries (due to trauma or tumour removal)

are still limited by the paucity of autologous materials available and donor site

morbidity. Recent advances in the development of scaffold-based Tissue Engineering

(TE) have given the surgeon new options for restoring form and function. There are

now bioactive biomaterials (second generation) available that elicit a controlled

action and reaction to the host tissue environment with a controlled chemical

breakdown and resorption to ultimately be replaced by regenerating tissue. Third-

generation biomaterials are now being designed to stimulate regeneration of living

tissues using tissue engineering and in situ tissue regeneration methods. Engineering

functional bone using combinations of cells, scaffolds and bioactive factors are seen

as a promising approach and these techniques will undoubtedly lead to ceaseless

possibilities for tissue regeneration and repair. There are currently thousands of

research papers and reviews available on bone tissue engineering, but there is still a

major discrepancy between scientific research efforts on bone tissue engineering and

the clinical application of such strategies. There is an evident lack of comprehensive

reviews that cover both the scientific research aspect as well as the clinical

translation and practical application of bone tissue engineering techniques. This

review will therefore discuss the state of the art of scientific bone tissue engineering

concepts and will also provide current approaches and future perspectives for the

clinical application of bone tissue engineering.


2.1 BONE BIOLOGY

Bone as an organ has next to its complex cellular composition a highly

specialised organic-inorganic architecture which can be classified as micro- and

nanocomposite tissue. Its mineralised matrix consists of (1) an organic phase (mainly

collagen, 35% dry weight) responsible for its rigidity, viscoelasticity and toughness;

(2) a mineral phase of carbonated apatite (65% dry weight) for structural

reinforcement, stiffness and mineral homeostasis; and (3) other non-collagenous

proteins that form a microenvironment stimulatory to cellular functions [83]. Bone

tissue exhibits a distinct hierarchical structural organization of its constituents on

numerous levels including macrostructure (cancellous and cortical bone),

microstructure (Harversian systems, osteons, single trabeculae), sub-microstructure

(lamellae), nanostructure (fibrillar collagen and embedded minerals) and sub-

nanostructure (molecular structure of constituent elements, such as mineral, collagen,

and non-collagenous organic proteins) (Figure 2) [84]. Macroscopically, bone

consists of a dense hard cylindrical shell of cortical bone along the shaft of the bone

that becomes thinner with greater distance from the centre of the shaft towards the

articular surfaces. Cortical bone encompasses increasing amounts of porous

trabecular bone (also called cancellous or spongy bone) at the proximal and distal

ends to optimise articular load transfer [83]. In humans, trabecular bone has a

porosity of 50-90% with an average trabecular spacing of around 1mm and an

average density of approximately 0.2 g/cm3 [85-87]. Cortical bone has a much denser

structure with a porosity of 3-12% and an average density of 1.80g/cm3

[86, 88].

On a microscopic scale, trabecular struts and dense cortical bone are composed

of mineralized collagen fibres stacked parallel to form layers, called lamellae (3–7

µm thick) and then stacked in a ± 45° manner [83]. In mature bone these lamellae

wrap in concentric layers (3–8 lamellae) around a central part named Haversian

canal which containings nerve and blood vessels to form what is called an Osteon (or

a Haversian system), a cylindrical structure running roughly parallel to the long axis

of the bone [84]. Cancellous bone consists of interconnecting framework of rod and

plate shaped trabeculae. On a nanostructural level, the most prominent structures are

the collagen fibres, surrounded and infiltrated by mineral. At the sub-nanostructural

level three main materials are bone crystals, collagen molecules, and non-

collagenous organic proteins. For further details the reader is referred to [84].


Figure 2: Hierarchical structural organization of bone (a) cortical and cancellous bone; (b) osteons

with Haversian systems; (c) lamellae; (d) collagen fibre assemblies of collagen fibrils; (e) bone

mineral crystals, collagen molecules, and non-collagenous proteins. Reproduced with permission from

(84), ©1998 IPEM.

There mineralised bone matrix is populated with four bone-active cells:

Osteoblasts, osteoclasts, osteocytes and bone lining cells. Additional cell types are

contained within the bone marrow that fills the central intramedullary canal of the

bone shaft and intertrabecular spaces near the articular surfaces [89]. Bone has to be

defined as an organ composed of different tissues and also serves as a mineral

deposit affected and utilised by the body’s endocrine system to regulate (among

others) calcium and phosphate homeostasis in the circulating body fluids.

Furthermore, recent studies indicate that bone exerts an endocrine function itself by

producing hormones that regulate phosphate and glucose homeostasis integrating the

skeleton in the global mineral and nutrient homeostasis [90].

Bone is a highly dynamic form of connective tissue which undergoes

continuous remodelling (the orchestrated removal of bone by osteoclasts followed by

the formation of new bone by osteoblasts) to optimally adapt its structure to changing

functional demands (mechanical loading, nutritional status etc.). From a material

science point of view bone matrix is a composite material of a polymer-ceramic

lamellar fibre-matrix and each of these design and material aspects influence the


mechanical properties of the bone tissue [91]. The mechanical properties depend on

the bone composition (porosity, mineralisation etc.) as well as the structural

organisation (trabecular or cortical bone architecture, collagen fibre orientation,

fatigue damage etc.) [92]. Collagen possesses a Young’s modulus of 1-2 GPa and an

ultimate tensile strength of 50-1000 MPa, compared to the mineral hydroxyapatite

which has a Young’s modulus of ~130GPa and an ultimate tensile strength of

~100MPa. The resulting mechanical properties of the two types of bone tissue,

namely the cortical bone and cancellous bone, are shown in Table 1. Age and related

changes in bone density have been reported to substantially influence the mechanical

properties of cancellous bone [93]. As outlined above, bone shows a distinct

hierarchical structural organization and it is therefore important to also define the

mechanical properties at microstructural levels (Table 2). Although the cancellous

and cortical bone may be of the same kind of material, the maturation of the cortical

bone material may alter the mechanical properties at the microstructural level.

Bone tissue is also known to be mechano-receptive; both normal bone

remodelling and fracture or defect healing are influenced by mechanical stimuli

applied at the regenerating defect site and surrounding bone tissue [94-97]. In

contrast to most other organs in the human body, bone tissue is capable of true

regeneration, i.e. healing without the formation of fibrotic scar tissue [98]. During

the healing process basic steps of fetal bone development are recapitulated and bone

regenerated in this way does not differ structurally or mechanically from the

surrounding undamaged bone tissue [99]. However, despite this tremendous

regenerative capacity, 5-10% of all fractures are prone to delayed bony union or will

progress towards a non-union and the development of a pseudarthrosis [100, 101].

Together with large traumatic bone defects and extensive loss of bone substance after

tumour resection or revision surgery after failed arthroplasties, these pathological

conditions still represent a major challenge in today’s clinical practice. The rangeof

bone graft materials available to treatsuch problems in modern clinical practice

essentially include autologous bone (from the same patient), allogeneic bone (from a

donor), and demineralised bone matrices, as well as a wide range of synthetic bone

substitute biomaterials such as metals, ceramics, polymers, and composite materials.

During the last decades, tissue engineering strategies to restore clinical function have

raised considerable scientific and commercial interest in the field of orthopaedic


surgery as well as reconstructive and oromaxillofacial surgery. Yet, the treatment of

bone defects and the search for bone substitute materials is not just a modern day

phenomenon, with its history reaching back through millennia.

Mechanical properties of compact and spongy bone[102]

Property Cortical bone Cancellous

bone

Compressive strength (MPa) 100-230 2-12

Flexural, tensile strength (MPa) 50-150 10-20

Strain to failure (%) 1-3 5-7

Fracture toughness (MPam1/2

) 2-12 -

Young’s modulus (GPa) 7-30 0.5-0.05

Table 1: Mechanical properties of compact (cortical) and spongy (cancellous) bone. Reproduced

and modified from (102).

Young’s modulus (GPa) (according to various levels of architecture)

Wet specimen (macrostructural)[103] 14-20

Wet specimen (microstructural)[104] 5.4

Dry specimen (submicrostructure)[105] 22

Table 2: Young’s modulus (GPa) (according to various levels of architecture). Modified from

(103-105) as listed in the table.

2.2 BONE GRAFTING AND BONE SUBSTITUTES IN THE LAST 4000

YEARS

The quest for the most efficient way to substitute for lost bone and to develop

the best bone replacement material has been pursued by humans for thousands of

years.

In Peru, archaeologists discovered the skull of a tribal chief from 2000 BC in

which a frontal bone defect (presumably from trepanation) had been covered with a

1mm-thick plate of hammered gold [106]. Trephined Incan skulls have been found

with plates made from shells, gourds, and silver or gold plates covering the defect

areas [107]. In a skull found in the ancient center of Ishtkunui (Armenia) from


approx. 2000 BC, a 7 mm diameter skull defect had been bridged with a piece of

animal bone [108]. These pursuits are not limited to skull surgeries involving bone

substitutes. Ancient Egyptians have been shown to have profound knowledge of

orthopaedic und traumatological procedures with Surgeons having implanted iron

prostheses for knee joint replacement as early as 600 BC, as analyses of preserved

human mummies have revealed [109].

The first modern era report of a bone xenograft procedure is believed to be the

Dutch surgeon Job Janszoon van Meekeren in 1668 [110, 111]. A skull defect of a

Russian nobleman was successfully treated with a bone xenograft taken from the

calvaria of a deceased dog. The xenograft was reported to have become fully

incorporated into the skull of the patient. In the 1800s, plaster of Paris (Calcium

sulphate) was used to fill bone cavities in patients suffering from Tuberculosis [112].

Attempts were also made to fill bone defects with cylinders made from ivory [113].

In 1820 the German surgeon Phillips von Walters described the first clinical use of a

bone autograft to reconstruct skull defects in patients after trepanation [114]. Walters

successfully repaired trepanation holes, following surgery to relieve intracranial

pressure, with pieces of bone taken from the patient’s own head. In 1881, Scottish

surgeon William MacEwen described the first allogenic bone grafting procedure: He

used tibial bone wedges from three donors that had undergone surgery for skeletal

deformity correction (caused by rickets) to reconstruct an infected humerus in a 3-

year-old child [115].

Major contributions leading to the development of modern day bone grafting

procedures and bone substitutes have been made by Ollier and Barth in the late

1800s. Louis Léopold Ollier carried out extensive experiments to study the

osteogenic properties of the periosteum and other various approaches to new bone

formation, mainly in rabbit and dog models. He also meticulously reviewed the

literature on bone regeneration available at that time and in 1867 he published his

1000-page textbook ‘Traite experimentel et clinique de la regeneration des os et de

la production artificielle du tissu osseux’, in which he described the term ‘bone graft’

(“greffe osseuse”) for the first time [116]. In 1895 the German surgeon Arthur Barth

published his treatise ‘Ueber histologische Befunde nach Knochenimplantationen’

(‘On histological findings after bone implantations’) presenting his results of various

bone grafting procedures involving the skull and long bones (humerus, forearm


bones) of dogs and rabbits including histological assessment [117]. Today, both

Ollier’s and Barth’s work are considered to be milestones in the development of

present day bone grafting procedures and bone substitute materials.

With the development of new orthopaedic techniques and increased numbers of

joint replacement procedures (prostheses), the demand for bone grafts increased in

the 20th

century, leading to the opening of the first bone bank for allogenic bone

grafts in New York in 1945 [118]. But the risk of an immunological reaction from

transplanted allogenic bone material was soon recognized and addressed in various

studies [119, 120]. Several procedures such as the use of hydrogen peroxide to

macerate bone grafts (“Kieler Span”) in the 1950s and 1960s to overcome antigenity

were not successful [121, 122]. Today, bone substitute materials such as (bovine)

bone chips are routinely used in clinical practice after being pre-treated to remove

antigen structures. However, due to the processing steps necessary to abolish

antigenicity, most of these grafts do not contain viable cells or growth factors and are

therefore inferior to viable autologous bone graft options. When allografts with

living cells are transplanted, there is a risk of transmitting viral and bacterial

infections: Transmission of human immunodeficiency virus (HIV), hepatitis C virus

(HCV), human T-lymphozytic virus (HTLV), unspecified hepatitis, tuberculosis and

other bacteria has been documented (mainly) for allografts (mainly from those

containing viable cells) [123].

As early as 1932, the work of the Swiss H. Matti proved the paramount

meaning of autologous cancellous bone grafts for bone regeneration approaches

[124]. Having conducted various experiments on the osteogenic potential of

autologous and allogenic bone, Schweiberer concluded in 1970 that the autologous

transplant remains the only really reliable transplantation material of the future, if

applied to bring about new bone formation or crucially to support the bridging bone

defects [125]. Even though this statement was made more than 50 years ago, it still

remains valid today, when bone is still the second most transplanted material, second

only to blood. Worldwide more than 3.5 million bone grafts (either autografts or

allografts) are performed each year [126]. Recent advances in technology and

surgical procedures have significantly increased the options for bone grafting

material, with novel products designed to replace both the structural properties of

bone, as well as promote faster integration and healing. The number of procedures


requiring bone substitutes is increasing, and will continue to do so as the population

ages and physical activity of the elderly population increases. Therefore, while the

current bone grafting market globally is estimated to be in excess of $2.5 billion US

each year, it is expected to increase at a compound annual growth rate of 7-8% [126].

Although the last decades have seen numerous innovations in bone substitute

materials, the treatment of bone defects with autologous bone grafting material is still

considered to be the ‘Gold Standard’ against which all other methods are compared

[47]. Autologous bone combines all the properties desired in a bone grafting

material: It provides a scaffold for the ingrowth of cells necessary for bone

regeneration (= osteoconductive); it promotes the proliferation of stem cells and their

differentiation into osteogenic cells (= osteoinductive) and it holds viable cells that

can form new bone tissue (= osteogenic) [99, 127]. However, the available volume of

autologous bone graft from a patient is limited and an additional surgical procedure

is required to harvest the grafting material which is associated with a significant risk

of donor site morbidity. 20-30% of autograft patients experience morbidity such as

chronic pain or dysaesthesia at the graft-harvesting site [51]. Large bone defects

(>5cm) may be treated with bone segment transport or free vascularized bone

transfer [58], as the use of an autologous bone graft alone is not recommended

because of the risk of graft resorption despite good soft tissue coverage [56]. The

vascularised fibula autograft [52] and the Ilizarov method [53-55] are the most

commonly used treatment methods for larger bone defects; however, complications

are common and the process can be laborious and painful for the patient as s/he may

be required to use external fixation systems for up to one and half years [57-59].

The limitations of existing bone grafting procedures, either autologous or

allogenic in nature, and the increased demand for bone grafts in limb salvage

surgeries for bone tumours and in revision surgeries of failed arthroplasties have

renewed the interest in bone substitute materials and alternative bone grafting

procedures [128]. In 1986, Masquelet and colleagues [129] first described a new

two-stage technique taking advantage of the body’s immune response to foreign

materials for bone reconstruction. The authors called it the ‘concept of induced

membranes’ – soon to become known as the ‘Masquelet technique’: In a first step, a

radical debridement of necrotic bone and soft tissue is followed by the filling of the

defect site with a polymethylmethacrylate (PMMA) spacer and stabilisation with an


external fixator. After the definitive healing of the soft tissue, a second procedure is

performed 6-8 weeks later, when the PMMA spacer is removed and a morsellised

cancellous bone graft (from the iliac crest) is inserted into the cavity [41, 42]. The

cement spacer was initially thought to prevent the collapse of the soft tissue into the

bone defect and to prepare the space for bone reconstruction. However, it was soon

discovered that the PMMA spacer does not only serve as a place holder, but that a

foreign body reaction to the spacer also induces the formation of a membrane that

possesses highly desirable properties for bone regeneration [42, 130]: The induced

membrane was shown to be richly vascularised in all layers; the inner membrane

layer (facing the cement) composed of synovial like epithelium and the out part is

made from fibroblasts, myoblasts and collagen. The induced membrane has also been

shown to secrete various growth factors in a time-dependent manner: High

concentrations of vascular endothelial growth factor (VEGF) as well as transforming

growth factor β (TGF β) are secreted as early as the second week after implantation

of the PMMA spacer; bone morphogenetic protein 2 (BMP-2) concentration peaks at

the fourth week. The induced membrane stimulates the proliferation of bone marrow

cells and differentiation towards an osteoblastic lineage. Finally, clinical experience

has shown that the cancellous bone inside the induced membrane is not subject to

resorption by the body. Ever since its introduction the ‘induced membrane’-

technique has been used very successfully in various clinical cases (see [41] and

references therein). However, the Masquelet technique still requires the harvesting of

an autologous bone graft, and with that come all the potential aforementioned

complications. Furthermore, the use of alternate bone substitute materials, such as

hydroxyapatite tricalcium phosphate, in combination with the Masquelet technique

has so far yielded results inferior to the use the Masquelet technique with autologous

bone grafting material [41, 131].

Besides the Masquelet technique, a more recent innovation has also

significantly improved the clinical approach to restoring bone defects. The

development of the Reamer-Irrigator-Aspirator (RIA©

)-System (DePuySynthes) has

given clinicians an alternative to iliac crest harvesting to retrieve bone grafting

materials from patients: The RIA System provides irrigation and aspiration during

intramedullay reaming, allowing the harvesting of finely morselised autologous bone

and bone marrow for surgical procedures requiring bone grafting material [132]. The


RIA was initially developed to lower the intramedullary pressure during the reaming

of long bones to reduce the risk of fat embolisms and pulmonary complications such

as the Acute Respiratory Distress Syndrom (ARDS), as well as to reduce local

thermal necrosis of bone tissue [133, 134]. However, the finely morsellised

autologous bone and bone marrow that is collected by the RIA has been shown to be

rich in stem cells, osteogenic cells and growth factors and has been recognized to be

a suitable bone graft alternative to the iliac crest autograft tissue [135, 136]. Also,

RIA enables the harvesting of larger bone graft volumes compared to the iliac crest

(approx. 40cm3 for the femur and 33cm

3 for the tibia) [51, 134]. Furthermore, the

risk of complications from the harvesting procedure has been reduced significantly

(RIA 6% vs. 19,37% for iliac crest autografts) [137]. Since its introduction, the

indications for use of RIA have been further extended to include the treatment of

postoperative osteomyelitis [138] and the harvesting of mesenchymal stem cells

(MSCs)[139]. The innovation driven by the RIA systems was so significant, that the

Journal “Injury” has dedicated a complete issue to the data available on RIA and its

applications recently [140]. A systematic review on the Reamer-irrigator-aspirator

indications and clinical results has recently been published by Cox et al. [141]. The

Masquelet technique as well as the RIA-system are nowadays frequently used in

clinical practice, independently. However, the two techniques may also be combined

to further improve their effectiveness when treating severe bone defects, for example

in post-traumatic limb reconstruction [142]. An example of a clinical case combining

the use of Masquelet technique and the use of the RIA-system to treat a complex case

of tibial non-union is provided in Figure 3.

Both the Masquelet technique and the development of the RIA-system

represent significant improvements in today’s clinical approach to bone

reconstruction and regeneration. However, utilising these techniques, we have still

not been able to replace autologous bone grafting in order to avoid surgical graft

retrieval procedures with all the associated disadvantages. However, with research

looking towards increasingly sophisticated bone tissue engineering techniques and

their first clinical applications the quest for developing improved bone substitute

material advances to the next level.


Figure 3: Clinical case combining the Masquelet-technique and the RIA-system to treat a tibial

non-union. 51 year old male acquired a Gustillo 3B fracture of the right tibia and fibula and was

treated with a stage procedure with locked plating and a free flap . The patient’s progress was very

slow and an implant failure occurred 8 months post-operatively (A). The patient was then referred for

the further management and underwent debridement of the non-union site on the distal tibia by lifting

the flap (B). The size of the extensive bone defect is shown in B (intraoperative image of situs and X-

ray image with retractor in defect site). Additionally, a PMMA bone cement spacer was inserted into

the tibial defect as part of the Masquelet technique. Postop X-ray images after surgery with the

PMMA spacer (circles) in place (C). 8 weeks later the PMMA spacer was removed and the induced

membrane at the defect site was packed with autologous cancellous bone graft obtained from the

femur using the Reamer-Irrigator-Aspirator (RIA) technique. (D) shows assembled RIA system, insert

showing morselised autologous bone and bone marrow graft obtained. Postop films after the second

surgery (E). 7 weeks after bone grafting the defect showed good healing and patient was able to fully

bear weight as tolerated. Over the following 2 months X-ray images showed progressive bridging of

the zone and he was able to return to work with light duties. He was reviewed again 7 months post-

surgery and had returned to work full-time and was walking long distances without any support (F)


2.3 BONE SUBSTITUTE MATERIALS (BSM)

Bone substitutes can be defined as “a synthetic, inorganic or biologically

organic combination - biomaterial - which can be inserted for the treatment of a bone

defect instead of autogenous or allogenous bone” [60]. This definition applies to

numerous substances and a variety of materials have been used over time in

anattempt to substitute bone tissue. Although merely of historic interest and with no

significance in modern therapies, the use of seashells, nuts, gourds and so forth show

that humans have strived for BSM for thousands of years.

With the introduction of tissue engineering and its clinical application the

regenerative medicine in 1993 [143] the modern day quest for BSMs has undergone

a significant change. The limitations of current clinical approaches have necessitated

the development of alternative bone repair techniques and have driven the

development of scaffold-based tissue engineering strategies. In the past, mostly inert

bone substitute materials have been used, functioning mainly as space holders during

the healing processes. Now a paradigm shift has taken place towards the use of new

‘intelligent’ tissue engineering biomaterials that would support and even promote

tissue re-growth [144].

According to the “diamond concept” of bone tissue engineering [145, 146] an

ideal bone substitute material should offer an osteoinductive three-dimensional

structure, contain osteogenic cells and osteoinductive factors, have sufficient

mechanical properties and promote vascularisation. Despite extensive research in the

field of bone tissue engineering, apart from the “gold standard” autograft bone, no

currently available BSM can offer these properties in one single material. Therefore,

the fundamental concept underlying tissue engineering is to combine a scaffold or

three-dimensional construct with living cells, and/or biologically active molecules to

form a “tissue engineering construct” (TEC), which promotes the repair and/or

regeneration of tissues [147, 148].

Currently used BSM can be classified into different subgroups according to

their origin [144, 149]:

1. BSM of natural origin

This group consists of harvested autogenous bone grafts as well as allogenic

BSM, such demineralised bone matrix, corticocancellous or cortical grafts,


cancellous chips (from either cadavers or living donors) [150-152]. Xenogenic

materials, for example porous natural bone hydroxyapatite from animal bones

(bovine, equine, porcine etc.) are also part of this group [153]. Phytogenic materials

such as bone-analogue calcium phosphate originally obtained from marine algae or

coral derived materials, also fall into this category [154, 155].

2. Synthetical (alloplastic) materials

This groups contains ceramics such as bioactive glasses [156],

Tricalciumphosphates (TCP) [157, 158], Hydroxyapatite (HA) [159-161] and glass

ionomer cements as well as Calcium Phosphate (CP) ceramics [162]. Metals such as

titanium also belong to this group. Furthermore polymers including

polymethylmethacrylate (PMMA), polylactides/poliglycolides and copolymers as

well as polycaprolactone (PCL)[163] are summarised in this group [144, 147, 164,

165].

3. Composite materials

BSM combining different materials such as ceramics and polymers are referred

to as composite materials [160, 166, 167]. By merging materials with different

structural and biochemical properties into composite materials, the properties of

composite materials can be modified to achieve more favourable characteristics, for

instance with respect to biodegradability [147, 165].

4. BSM combined with growth factors

Natural or recombinant growth factors such a bone morphogenic protein

(BMP), platelet-derived growth factor (PDGF), transforming growth factor-ß (TGF-

β), insulin-like growth-factor 1, vascular endothelial growth factor (VEGF) and

fibroblast growth factor can be added to increase the biological activity of BSM

[168, 169]. For example, a composite material made of medical-grade

polycaprolactone-tricalcium phosphate (mPCL-TCP) scaffolds (combined with

recombinant human BMP-7) has been demonstrated to completely bridge a critical-

sized (3cm) tibial defect in a sheep model [75].

5. BSM with living cells

Mesenchymal stem cells [170-172], bone marrow stromal cells [173, 174],

periosteal cells [175, 176], osteoblasts [177] and embryonic [178] as well as adult


stem cells [179] have been used in bone tissue engineering [80, 99, 169, 180-182].

These cells can generate new tissue alone or can be used in combination with

scaffold matrices.

BSMs can also be classified according to their properties of action. An

overview of the currently available BSM for clinical (orthopaedic) use and their

mode of action is given in Table 3 (reproduced from [183]).

2.4 THREE-DIMENSIONAL SCAFFOLDS IN BONE TISSUE

ENGINEERING

Scaffolds serve as three-dimensional structures to guide cell migration,

proliferation and differentiation. In load bearing tissues, it also serves as temporary

mechanical support structure. Scaffolds substitute for the function of the extracellular

matrix and need to fulfil highly specific criteria. An ideal scaffold should be (i) three-

dimensional and highly porous with an interconnected pore network for cell growth

and flow transport of nutrients and metabolic waste; (ii) should have surface

properties which are optimized for the attachment, migration, proliferation and

Table 3: Bone grafts and graft substitutes currently used in clinical orthopaedic applications.

Reproduced with permission from (183), © The IJMR, 2010.


differentiation of cell types of interest (depending on the targeted tissue); (iii) be

biocompatible, not elicit an immune response and be biodegradable with a

controllable degradation rate to compliment cell/tissue in-growth and maturation; (iv)

its mechanical properties should match those of the tissue at the site of implantation

and (v) the scaffold structure should be easily and efficiently reproducible in various

shapes and sizes [165].

2.4.1 Biocompatibility

Biocompatibility represents the ability of a material to perform with an

appropriate response in a specific application [184]. As a general rule, scaffolds

should be fabricated from materials that do not have the potential to elicit

immunological or clinically detectable primary or secondary foreign body reactions

[185]. Parallel to the formation of new tissue in vivo, the scaffold may undergo

degradation via the release of by-products that are either biocompatible without proof

of elimination form the body (biodegradable scaffolds) or can be eliminated through

natural pathways from the body, either by simple filtration of by-products or after

their metabolisation (bioresorbable scaffolds) [165]. Due to poor vascularisation or

low metabolic activity, the capacity of the surrounding tissue to eliminate the by-

products may be low leading to a build up of the by-products thereby causing local

temporary disturbances [165]: A massive in vivo release of acidic degradation by-

products leading to inflammatory reactions has been reported for several

bioresorbable devices made from polylactides [186-188]. Another example is the

increase of osmotic pressure or pH caused by local fluid accumulation or transient

sinus formation from fibre reinforced polyglycolide pins used in orthopaedic

applications [186]. It is also known that calcium phosphate biomaterial particles can

cause inflammatory reactions after being implanted (although this inflammatory

reaction may be considered desirable to a certain extent as it subsequently stimulates

osteoprogenitor cell differentiation and bone matrix deposition) [189]. These

examples illustrate that potential problems related to biocompatibility in tissue

engineering constructs for bone and cartilage applications may be related to the use

of biodegradable, erodible and bioresorbable polymer scaffolds. Therefore, it is

important that the three dimensional Tissue Engineering Construct (TEC) is exposed

at all times to sufficient quantities of neutral culture media when undertaking cell

culture procedures, especially during the period where the mass loss of the polymer


matrix occurs [165]. For applications in vivo, it is of course not possible to expose

the TEC to neutral media, and one therefore has to carefully take into account the

local specifications (pH, vascularisation, metabolic activity etc) of the tissue to be

engineered when accessing biocompatibility of a TEC.

2.4.2 Mechanical properties and degradation kinetics

The design of tissue engineering scaffolds needs to consider physio-chemical

properties, morphology and biomechanical properties as well as degradation kinetics.

The scaffold structure is expected to guide the development of new bone formation

by promoting attachment, migration, proliferation and differentiation of bone cells.

Parallel to tissue formation, the scaffold should also undergo degradation in order to

allow for ultimate replacement of scaffold material with newly formed, tissue

engineered bone. Furthermore, the scaffold is also responsible for (temporal)

mechanical support and stability at the tissue engineering site until the new bone is

fully matured and is able to withstand mechanical load. As a general rule, the

scaffold material should be sufficiently robust to resist changes in shape resulting

from the introduction of cells into the scaffold (each of which should capable of

exerting tractional forces) and from wound contraction forces that would be evoked

during tissue healing in vivo [147]. In order to achieve optimal results, it is therefore

necessary to carefully balance the biomechanical properties of a scaffold with its

degradation kinetics. A scaffold material has to be chosen that degrades and resorbs

at a controlled rate, giving the TEC sufficient mechanical stability at all times, but at

the same time allowing new in vivo formed bone tissue to substitute for its structure.

Figure 4 depicts the interdependence of molecular weight loss and mass loss of a

slow degrading composite scaffold and also shows the corresponding stages of tissue

regeneration [148].

At the time of implantation the biomechanical properties of a scaffold should

match the structural properties of the tissue it is implanted into as closely as possible

[190]. It should possess sufficient structural integrity for the period until the

engineered tissue ingrowth has replaced the slowly disappearing scaffold matrix with

regards to mechanical properties. In bone tissue engineering the degradation and


resorption kinetics of the scaffold have to be controlled in such a way that the

bioresorbable scaffold retains its physical properties for at least 6 months to enable

Figure 4: Schematic illustrating the interdependence of molecular weight loss and mass loss of a slow-

degrading composite scaffold plotted against time, which corresponds with tissue regeneration. Scaffold, as

shown by SEM (a) is implanted at t = 0 (b) with lower figures (c-e) showing a conceptual illustration of the

biological processes of bone formation over time. The scaffold is immediate filled with a hematoma on

implantation (c) followed by vascularization (d) and gradually new bone is formed within the scaffold (e). As the

scaffold degrades over time there is increased bone remodeling within the implant site until eventually the

scaffold pores are entirely filled with functional bone and vascularity. SEM of scaffold degraded over time (g)

with associated schematic visualization of how mPCL-TCP scaffolds degrade via long-term bioerosion process,

which takes up to 36 months in vivo (h). Reproduced with permission from (148), © Elsevier Ltd 2012


cell and tissue remodelling to achieve stable biomechanical conditions and

vascularisation at the defect site [165]. Apart from host anatomy and physiology, the

type of tissue that is aimed to be engineered also has a profound influence on the

degree of remodelling: in cancellous bone the remodelling takes 3-6 months, while

cortical bone will take twice as long, approximately 6-12 months, to remodel [147].

Whether the TEC will be part of a load bearing or non-load bearing site will also

significantly influence the needs for mechanical stability of the TEC as mechanical

loading can directly affect the degradation behaviour as well [147]. Utilising

orthopaedic implants to temporarily stabilise the defect area also influences the

requirements for biomechanical stability of the TEC significantly [95, 191]. It is

therefore crucial to meticulously select the scaffold material individually for each

tissue engineering approach to tailor the mechanical properties and degradation

kinetics exactly to the purpose of the specific TEC [165]. Consequently, there is not

one “ideal scaffold material” for all bone tissue engineering purposes, but the choice

depends on the size, type and location of the bone tissue to be regenerated.

2.4.3 Surface Properties

The surface area of a scaffold represents the space where pivotal interactions

between biomaterial and host tissue take place. The performance of a TEC depends

fundamentally on the interaction between biological fluids and the surface of the

TEC, and it is often mediated by proteins absorbed from the biological fluid [192].

The initial events include the orientated adsorption of molecules from the

surrounding fluid, creating a specific interface to which the cells and other factors

respond to the macrostructure of the scaffold as well as the microtopography and

chemical properties of the surface determine which molecules are adsorbed and how

cells will attach and align themselves [193]. The focal attachments made by the cells

with their substrate then determines cell shape, which in turn transduces signals via

the cytoskeleton to the nucleus resulting in expression of specific proteins which may

be structural or signal-related and contribute towards the cell phenotype.

Due to technical progress, we are now able to manipulate materials at the

atomic, molecular, and supramolecular level, and bulk materials and surfaces can be

designed at a similar dimension to that of the nanometer constituent components of


bone [83]: In natural bone, hydroxyapatite plates are approximately between 25nm in

width and 35nm in length while collagen type 1 is a triple helix 300nm in length,

0.5nm in width and with a periodicity of 67nm [194]. “Nanomaterials” commonly

refers to materials with basic structural units in the range 1–100 nm (nanostructured),

crystalline solids with grain sizes between 1 and 100 nm (nanocrystals), individual

layers or multilayer surface coatings in the range 1–100 nm (nanocoatings),

extremely fine powders with an average particle size in the range 1–100 nm and

fibres with a diameter in the range 1–100 nm (nanofibres) [83]. The close proximity

of the scale of these materials to the scale of natural bone composites makes the

application of nanomaterials for bone tissue engineering a very promising strategy.

Surfaces with nanometer topography can promote the availability of amino acid and

proteins for cell adhesion to a great extent, for example, the adsorption of fibronectin

and vitronectin (two proteins known to enhance osteoblast and bone forming cell

function [195]) can be significantly increased by decreasing the grain size on the

scaffold/implant surface below 100nm [196]. It has also been shown that calcium-

mediated cell protein adsorption on nanophase material promotes unfolding of these

proteins promoting bone cell adhesion and function [196]. Current literature supports

the hypothesis that by creating surface topographies with characteristics that

approximate the size of proteins, a certain control over protein adsorption and

interactions will be possible. Since the surface characteristics regarding of roughness,

topography and surface chemistry are then transcribed via the protein layer into

information that is comprehensible for the cells [193], this will enable the fabrication

of surface properties directly targeted at binding specific cell types. In vitro,

osteoblast adhesion, proliferation and differentiation and calcium deposition is

enhanced on nanomaterials with grain sizes less than 100nm [196, 197]. The

adherence of osteoblasts has been shown to increase up to threefold when the surface

is covered with nanophase titanium particles instead of conventional titanium

particles [198]. Nano- and microporosity has also been shown to promote osteogenic

differentiation [199] and osteogenesis [200]. The use of nanomaterials to achieve

better osteointegration of orthopaedic implants and for bone tissue engineering

approaches has been extensively summarised in several recent reviews [83, 201-204]

and will not be reviewed in its entirety here.


However, it becomes clear that rough scaffold surfaces favour attachment,

proliferation and differentiation of anchorage-dependent bone forming cells [205].

Osteogenic cells migrate to the scaffold surface through a fibrin clot initially

established immediately after implantation of the TEC from the haematoma caused

by the surgical procedure [169]. The migration causes retraction of the temporary

fibrin matrix and, if not well secured, can lead to detachment of the fibrin from the

scaffold during wound contraction leading to decreased migration of the osteogenic

cells into the scaffold [206, 207].

With regards to surface chemistry, degradation properties and by-products

(relating to pH, osmotic pressure, inflammatory reactions etc.) are of importance and

have been briefly discussed already. In the following section, the role of calcium

phosphate in the osteoinductivity of biomaterials will be summarized as an example

of how surface chemistry may be manipulated to benefit scaffold properties. To date,

most synthetic biomaterials that have been shown to be osteoinductive contained

calcium phosphate underlining the crucial role of calcium and phosphate in

osteoinduction properties of biomaterials [208]. As summarised above, adequate

porosity and pore size is crucial for bone tissue engineering scaffolds in order to

allow sufficient vascularisation and enable a supply of body fluids throughout the

TEC. Together with this nutrient supply, a release of calcium and phosphate ions

from the biomaterial surface takes places and is believed to be the origin of

bioactivity of calcium phosphate biomaterials [209-211]. This process is followed by

the precipitation of a biological carbonated apatite layer (that contains calcium-,

phosphate- and other ions such as magnesium as well as proteins and other organic

compounds) that occurs when the concentration of calcium and phosphate ions has

reached supersaturation level in the vicinity of the implant [208, 212, 213]. This

bone-like biological carbonated apatite layer is thought to be physiological trigger for

stem cells to differentiate down the osteogenic lineage or could induce the release of

growth factors that complement this process [208]. For biomaterials lacking calcium

phosphate particles, the roughness of the surface is considered to act as a collection

of nucleation sites for calcium phosphate precipitation from the hosts’ body fluids,

thereby forming a carbonated apatite layer.

Comparing calcium phosphate (CaP) coated fibrous scaffolds (fibre diameter

approx 50um) made from medical grade polycaprolactone (mPCL) with non- coated


mPCL-scaffolds, we have shown that CaP-coating is beneficial for new bone

formation in vitro, enhancing alkaline phosphatase activity and mineralisation within

the scaffolds [214]. Interestingly, other research has shown that the implantation of

highly soluble carbonated apatite ceramics alone did not result in bone induction in

vivo [215], suggesting that a relatively stable surface (e.g. through a composite

material that contains a less soluble phase) is needed for the facilitation of bone

formation as discussed above (see “mechanical properties and degradation kinetics”).

Bone formation requires a stable biomaterial interface and therefore, too rapid in vivo

dissolution of calcium phosphate materials has been shown to be unfavourable for

the formation of new bone tissue [216, 217]. Chai et al. and Barradas et al. have

recently reviewed the effects of calcium phosphate osteogenicity in bone tissue

engineering [216, 218].

Further comprehensive reviews on the influence of surface topography and

surface chemistry on cell attachment and proliferation for orthopaedic implants and

bone tissue engineering are available [83, 192, 208, 216, 219].

2.4.4 Porosity and pore size

Porosity is commonly defined as the percentage of void space in a so called

cellular solid (the scaffold in bone tissue engineering applications) [220]. Using solid

and porous particles of hydroxyapatite for the delivery of the growth factor BMP-2,

Kuboki et al showed that pores are crucial for bone tissue formation because they

allow migration and proliferation of osteoblasts and mesenchymal cells, as well as

vascularisation; no new bone formed on solid particles [221]. A porous scaffold

surface also improves mechanical interlocking between the implanted TECs and the

surrounding natural bone tissue, providing greater mechanical stability at this crucial

interface in tissue engineering [222].

Scaffold porosity and pore size relate to the surface area available for the

adhesion and growth of cells both in vitro as well as in vivo and to the potential for

host tissue ingrowth, including vasculature, to penetrate into the central regions of

the scaffold architecture. In assessing the significance of porosity several in vivo

studies have been conducted utilising hard scaffold materials such as calcium

phosphate or titanium with defined porous characteristics [223]. The majority of


these studies indicate the importance of pore structure in facilitating bone growth.

Increase of porosity as well as pore size and spacing of pore interconnectivity has

been found to positively influence bone formation in vivo, which is also correlated

with scaffold surface area. Pore interconnections smaller than 100μm were found to

restrict vascular penetration and supplementation of a porous structure with

macroscopic channels has been found to further enhance tissue penetration and bone

formation [165, 224]. Interestingly, these results correlate well with the diameter of

the physiological Haversian systems in bone tissue that possess an approximate

diameter of more than 100µm. The ability of new capillary blood vessels to grow

into the TEC is also related to the pore size, thereby directly influencing the rate of

ingrowth of newly formed bone tissue into the TEC: In vivo, larger pore sizes and

higher porosity lead to a faster rate of neovascularisation, thereby promoting greater

amounts of new bone formation via direct osteogenesis. In contrast, small pores

favour hypoxic conditions and induce osteochondral formation before osteogenesis

occurs [160]. Pores and pore interconnections should be at least 300 microns in

diameter to allow sufficient vascularisation. Besides the actual macroporosity (pore

size >50µm) of the scaffold microporosity (pore size <10µm) and pore wall

roughness also have a large impact on osteogenic response: Microporosity results in

larger surface areas contributing to higher bone-inducing protein adsorption and to

ion exchange and bone-like apatite formation by dissolution and re-precipitation

[205, 223]. As outlined above, sub-micron and nanometre surface roughness favours

attachment, proliferation and differentiation of anchorage-dependent bone forming

cells [205].

Although increased porosity and higher pore size facilitate bone ingrowth, it

also compromises the structural integrity of the scaffold, and if the porosity becomes

too high it may adversely affect the mechanical properties of the scaffold at the same

time [147]. In addition, the rate of degradation is influenced by the porosity and pore

size (for biodegradable scaffolds). A higher pore surface area enhances interaction of

the scaffold materials with host tissue and can thereby accelerate degradation by

macrophages via oxidation and/or hydrolysis [223]. Therefore, scaffolds fabricated

from biomaterials with a high degradation rate should not have high porosities

(>90%) in order to avoid compromise to the mechanical and structural integrity

before adequate substitution by newly formed bone tissue. Scaffolds made from


slowly degrading biomaterials with robust mechanical properties can, in contrast, be

highly porous [223]. Table 4 [147] illustrates mechanical properties and degradation

kinetics in relation to the porosity for many commonly used composite scaffolds.

This illustrates that there are a number of advantages and disadvantages associated

with any changes made to the porosity or pore size of scaffolds. It is inevitable to

find a balance between these pros and cons in order to tailor the scaffold properties

ideally to the demands of the tissue engineering approach used. For comprehensive

reviews on role of porosity and pore size in tissue engineering scaffolds, the reader is

referred to two recently published reviews [223, 225].

It becomes clear that a multitude of factors has to be taken into account when

designing and fabricating scaffolds for bone tissue engineering. However, it is

beyond the scope of this review to present all of them in detail and a number of

comprehensive reviews have been published recently on this topic [83, 86, 147, 165,

169, 226, 227].

Table 4: Mechanical properties and degradation kinetics in relation for porosity of composite

scaffolds. Reproduced with permission from (147), Copyright © 2007 John Wiley & Sons, Ltd.


2.5 ADDITIVE MANUFACTURING AND COMPUTER AIDED DESIGN –

GAME CHANGERS IN THE FABRICATION OF THREE-

DIMENSIONAL SCAFFOLDS

The three-dimensional design characteristics in combination with the material

properties of a scaffold are crucial for bone tissue engineering purposes. Not only

does the scaffold structure need to be controlled on a macroscopic level (to achieve

sufficient interposition of the scaffold into the defect site), but also on a microscopic

level (to optimise tissue engineering properties with regards to osteoinduction,

osteoconduction, osteogenesis and vascularisation as well as mechanical stability)

and even down to nanostructural configuration (to optimise protein adsorption, cell

adhesion, differentiation and proliferation related to desired tissue engineering

characteristics of the TEC). It is therefore necessary to exert strict control over the

scaffold properties during the fabrication process. Conventional techniques for

scaffold fabrication include solvent casting and particulate leaching, gas foaming,

fibre meshes and fibre bonding, phase separation, melt molding, emulsion freeze

drying, solution casting and freeze drying [228]. All of these techniques are

subtractive in nature, meaning that parts of the fabricated scaffold are removed from

the construct after the initial fabrication process in order to generate the desired

three-dimensional characteristics. Hence a number of limitations exist regarding

these fabrication methods: conventional methods do not allow a precise control over

pore size, pore geometry, pore interconnectivity or spatial distribution of pores and

interconnecting channels of the scaffolds fabricated [160, 229, 230]. In addition,

many of these techniques require the application of organic solvents and their

residues can impose severe adverse effects on cells due to their potentially toxic

and/or carcinogenic nature, reducing the biocompatibility of the scaffold

significantly [231].

The introduction of additive manufacturing (AM) techniques into the field of

bone tissue engineering has helped to overcome many of these restrictions [160, 228,

232]. In AM three-dimensional objects are created in a computer-controlled layer-by-

layer fabrication process. In contrast to subtractive conventional methods of scaffold

fabrication, this technique is additive in nature and does not involve removal of

materials after the initial fabrication step. These techniques have also been named

“rapid prototyping” or “solid free form fabrication” in the past, but in order to clearly

distinguish them from conventional methods the latest ASTM standard now


summarises all of these techniques under the term “Additive Manufacturing” [233].

The basis for each AM process is the design of a three-dimensional digital or in silico

model of the scaffold to be produced. This computer model can either be created

from scratch using “computer aided design” (CAD) methods or can be generated

using data from a 3D-scan of existing three-dimensional structures (such as the

human skeleton) [234]. The digital model is then converted into an STL-file that

expresses the three-dimensional structure as the summary of multiple horizontal two-

dimensional planes. Using this STL-file an AM-machine then creates the three-

dimensional scaffold structure in a layer-by-layer fabrication method in which each

layer is tightly connected to the previous layer to create a solid object. A number of

different AM techniques are currently applied using thermal, chemical, mechanical

and/or optical processes to create the solid three-dimensional object [232]. These

methods include laser-based methods such as Stereolithography (STL) and Selective

Laser Sintering (SLS), printing-based applications (e.g. 3D-Printing, Wax-Printing)

and Nozzle-based systems like Melt Extrusion/Fused Deposition Modeling (FDM)

and Bioplotting. The multitude of AM techniques and their specifications were

reviewed by several authors lately [228, 232, 235, 236].

AM techniques have been used since the 1980s in the telecommunication

industry, in jewellery making and production of automobiles [237]. From the 1990s

onwards, AM was gradually introduced to the medical field as well [238]: AM was

initially used to fabricate three-dimensional models of bone pathologies in

orthopaedic maxillofacial neurosurgical applications to plan surgical procedures and

for haptic assessment during the surgery itself [239, 240]. With recent technical

advances AM is nowadays applied to make custom-made implants and surgical tools

[241] and to fabricate highly detailed, custom-made threedimensional models for the

individual patient (using data from CT, MRI, SPECT etc.) to plan surgical

approaches, specifically locate osteotomy sites, choose the correct implant and to

predict functional and cosmetic outcomes of surgeries [242, 243]. Thereby the

operating time as well as the risk of complications has been reduced significantly.

The application of AM in bone tissue engineering represents a highly

significant innovation that has drastically changes the way scaffolds are being

fabricated; AM has more or less become the new gold standard for scaffold

manufacturing [160]. The advantages of rapid prototyping processes include (but are


not limited to) increased speed, customisation and efficiency. AM technologies have

relatively few process steps and involve little manual interaction, therefore, three-

dimensional parts can be manufactured in hours and days instead of weeks and

months. The direct nature of AM allows the economical production of customized

tissue engineering scaffolds. The products can be tailored to match the patient’s

needs and still sustain economic viability as compared to traditional techniques

which must manufacture great numbers of devices. The conventional scaffold

fabrication methods commonly limit the ability to form complex geometries and

internal features. AM methods reduce the design constraints and enable the

fabrication of desired delicate features both inside and outside the scaffold. Using

STL, the AM technique with the highest precision, for example objects at a scale of

20µm can be fabricated [244]. A two-photon STL-technique to initiate the

polymerisation can be used to produce structures even at micrometer and sub-

micrometer levels [245].

AM methods allow for variation of composition of two or more materials

across the surface, interface, or bulk of the scaffold during the manufacturing.

Thereby, positional variations in physicochemical properties and surface

characteristics can be created and utilized to promote locally specific tissue

engineering signals. Several AM techniques operate without the use of toxic organic

solvents. This is a significant benefit, since incomplete removal of solvents may lead

to harmful residues that can affect adherence of cells, activity of incorporated

biological agents or surrounding tissues as already described. AM allows the control

of scaffold porosity leading to the applications that may have areas of greater or

lesser structural integrity and areas of encouraged blood flow due to increased

porosity. Fabricating devices and/or implants with differences in spatial distribution

of porosities, pore sizes, mechanical and chemical properties can mimic the complex

composition and architecture of natural bone tissue and thereby optimise bone tissue

engineering techniques. In addition, scaffolds with gradients in porosity and pore

sizes can be functionalised to allow vascularisation and direct osteogenesis in one

area of the scaffold, while promoting osteochondral ossification in the other, which is

an appealing approach to reproduce multiple tissues and tissue interfaces within one

and the same biomaterial scaffold [223].


Table 5 summarises the advantages of scaffolds designed and fabricated by AM

techniques.

Advantages of scaffolds designed and fabricated via additive manufacturing

Higher variability of designing a targeted degradability and resorbility as well as

improved biocompatiblity

Can be processed into various shapes, volumes and microstructures

Easily mass-produced or properties can be tailored for patient-specific applications

(addressing the scheme of Personalised Medicine)

Control over chemical and physically structural properties, crystallinity, hydrophobicity,

degradation rate and mechanical properties (e.g. through the alteration of surface

chemistry)

Allow exact engineering of matrix configuration, satisfying the biophysical limitations of

mass transfer

Flexibility to alter the physical properties and potentially facilitate reproducibility and

scale-up

Flexibility to manipulate the configuration of matrix to vary the surface area available for

cell attachments, also to optimize the exposure of attached cells to nutrients and allow

transport of waste products

The designs and fabrication of composite scaffolds which chemical environment

surrounding a synthetic degradable polymer material (e.g. aliphatic polyesters) be affected

in a controlled fashion as the polymer by-products are neutralized by ceramic

components

The potential to deliver continuously the nutrients and hormones that can be incorporated into the

scaffold structure

The ratio of surface area to mass can be altered or the porosity, pore size and pore size distribution

of the differing configurations can be altered so as to increase or decrease the mechanical

properties of the scaffold

Table 5: Advantages of scaffolds designed and fabricated via additive manufacturing


2.6 TRANSLATING BONE TISSUE ENGINEERING STRATEGIES FROM

BENCH TO BEDSIDE

Musculoskeletal conditions are highly prevalent and cause a large amount of

pain, illness and disability to patients. These conditions are the second most common

reason for consulting a general practitioner, accounting for almost 25% of the total

cost of illness and up to 15% of primary care [246]. In addition, the impact of

musculoskeletal conditions is predicted to grow with the increasing incidence of

lifestyle-related obesity, reduced physical fitness and increased road traffic accidents

[246]. The impact of bone trauma is significant – the consequences of failing to

restore full function to an injured limb are dramatically demonstrated by the statistic

that only 28% of patients suffering from severe open fractures of the tibia are able to

resume full function and hence return to previous employment [246]. Along with

trauma, tumour resection is another major cause of large bone defects. Cancer is a

major public health challenge, with one in four deaths in the United States currently

due to this disease. Recent statistics indicate that 1,638,910 new cancer cases and

577,190 deaths from cancer are projected to occur in the United States in 2012 [247].

As outlined above, the number of procedures requiring bone implant material is

increasing, and will continue to do so in our aging population and with deteriorating

physical activity levels [128]. The current bone grafting market already is estimated

to be in excess of $2.5 billion each year and is expected to increase by 7-8% per year

[126]. With the introduction of tissue engineering the hopes and expectations were

extremely high to be able to substitute natural organs with similar (or even better)

tissue engineered replacement organs. However, at the time it was stated that “few

areas of technology will require more interdisciplinary research than tissue

engineering” [143] and this assessment holds true today.

In the years to follow, numerous private and public institutes conducted

scientific research and clinical translation efforts related to tissue engineering. At the

beginning of 2001, tissue engineering research and development was being pursued

by 3,300 scientists and support staff in more than 70 start-up companies or business

units with a combined annual expenditure of over $600 million USD [248]. The US

National Institutes of Health (NIH), accounting for the largest cumulative US federal

research expenditures, has increased the funding in tissue engineering from 2.36

billion USD in the fiscal year 2003 to more than 614 billion USD for the fiscal year


2006 [249]. Between 2000 and 2008 the number of papers published on tissue

engineering and scaffolds per year increased by more than 400% and more than

900%, respectively [65]. But despite the increasing research expenditure and the

magnitude of discoveries and innovations in bone tissue engineering since its

introduction more than three decades ago, the translation of these novel techniques

into routine clinical applications on a large scale has still not taken place. As Scott J.

Hollister has pointed out, there is, on the one hand, a stark contrast between the

amount of tissue engineering research expenditures over the last 20 years and the

resulting numbers of products and sales figures. On the other hand, there is also a

significant discrepancy between the complexities of intended tissue engineering

therapies compared to the actual therapies that have reached clinical applications

[65]. This evident gap between research and clinical application/commercialisation is

commonly termed the “Valley of Death” due to the large number of ventures that

“die” between scientific technology development and actual commercialization due

to lack of funds (Figure 5) [65].

Figure 5: For tissue engineering, the Valley of Death is the gap and associated funding

difficulties of taking tissue engineering technologies to tissue-engineered products. The Valley

exists due to the need of obtaining funding to develop scalable/GMP design and manufacturing

processes, the need for pre-clinical studies proving therapies in large animal models, and finally, the

need to progress to clinical trials. Reproduced with permission from (65), © 2009 IOP Publishing Ltd.

All rights reserved.


The Valley of Death is particularly large for tissue engineering approaches

because this field of research often utilises immensely cost intensive high-tech

biotechnologies for technological development eating up large parts of the funding

available, but then additionally faces the challenges of funding large scale pre-

clinical studies and clinical studies to gain approval by regulatory bodies,

demonstrate product safety and gain clinical acceptance [65, 71, 250].

2.6.1 Bridging the gap between tissue engineering research and clinical

applications

To bridge the gap between the bench and bedside, the scaffold is required to

perform as a developmentally conducive extracellular niche, at a clinically relevant

scale and in concordance with strict clinical (economic and manufacturing)

prerequisites (Figure 6) [251]. In this context the scaffold facilitates for smaller and

medium sized defects the entrapment of the hematoma and prevents it’s “too early”

contraction [252]. For large and high-load bearing defects the scaffold can also

deliver cells and/or growth factors to the site of damage and provides an appropriate

template for new tissue formation. The scaffold should thus constitute a dynamically

long-lasting yet degradable three-dimensional architecture, preferably serving as a

functional tissue substitute which, over time, can be replaced by cell-derived tissue

function. Designing and manufacturing processes are believed to be the gatekeepers

to translate tissue engineering research into clinical tissue engineering applications

and concentration on the development of these entities will enable scaffolds to bridge

the gap between research and clinical practice [65]. One of the greatest difficulties in

bridging the Valley of Death is to develop good manufacturing processes and

scalable designs and to apply these in preclinical studies; for a description of the

rationale and road map of how our multidisciplinary research team has addressed this

first step to translate orthopaedic bone engineering from bench to bedside see below

and refer to our recent publication [71]. In order to take bone tissue engineering

approaches from bench to bedside, it also imperative to meticulously assess the

clinical demands for specific scaffold characteristics to achieve a broad and

optimised range of clinical applications for the specific tissue engineering approach.

A sophisticated bone tissue engineering technology will not necessarily have

multiple clinical applications just because of its level of complexity, and defining

specific clinical target applications remains one of the most underestimated


challenges in the bridging the Valley of Death [65]. There is often a great level of

discrepancy between the clinical demands on a tissue engineering technique and the

scientific realisation of such technique, hampering the clinical translation. Thus a

scaffold that is realistically targeted at bridging the Valley of Death should [251]: (i)

meet FDA approval (for further details on this topics see reviews by Scott J. Hollister

2011 and 2009) [65, 67]; (ii) allow for cost effective manufacturing processes; (iii)

be sterilisable by industrial techniques; (iv) enable easy handling without extensive

preparatory procedures in the operation theatre; (v) preferably, be radiographically

distinguishable from newly formed tissue; and (vi) allow minimally invasive

implantation [253, 254].

Figure 6: Bone tissue engineering strategies rely on three-dimensional scaffolds that constitute an

inductive/conductive extracellular microenvironment for stem cell function as well as a delivery

vehicle and 3D scaffold of clinically relevant properties and proportions. In fulfilling these dual

criteria the biomimetic scaffold plays a critical role bridging the gap between the developmental

context of stem cell mediated tissue formation and the adult context of injury and disease. Reproduced

with permission from (251), © 2008 Elsevier Inc.


2.6.2 Rationale for translating bone tissue engineering strategies into clinical

applications

In targeting the translation of a (bone) tissue engineering approach from bench

to bedside, there is a distinct hierarchy and sequence of the type of studies that need

to be undertaken to promote the translation process [255]: Having identified clinical

needs and based on fundamental discoveries regarding biological mechanisms, a

novel tissue engineering approach is designed and first studies are undertaken to

characterise mechanical and chemical properties of the TEC to be used. The next

step involves feasibility and bioactivity testing and should be carried out in vitro and

in vivo. In vitro assays using cell culture preparations are used to characterise the

effects of materials on isolated cell function and for screening large numbers of

compounds for biological activity, toxicity and immunogenicity [256, 257].

However, due to their nature using isolated cells, in vitro models are unavoidably

limited in their capacity to reflect complex in vivo environments that the TEC will be

exposed to and are therefore inadequate to predict in vivo or clinical performances.

Therefore, in vivo models (that is animal models) are required in order to overcome

the limitations of in vitro models to provide a reproducible approximation of the real

life situation. In vivo feasibility testing is almost exclusively done in small animals,

mainly in rodents and rabbits [255, 258-260]. The advantages of small animal

models include relatively easy standardisation of experimental conditions, fast bone

turnover rates (= shorter periods of observation), similar lamellar bone architecture

and similar cancellous bone thinning and fragility, similar remodelling rates and

sites, common availability and relatively low costs for housing and maintenance.

Disadvantages of rodent and rabbit models include different skeletal loading patterns,

open epiphyses at various growth plates up to the age of 12-14 months (or for

lifetime in rats), minimal intra-cortical remodelling, the lack of Harversian canal

systems, a smaller proportion of cancellous bone to total bone mass and their

relatively small size for testing of implants [259]. Whilst a large number of studies in

rodents and rabbits have established proof of concept for bone tissue engineering

strategies, scaling up to larger, more clinically relevant animal models has presented

new challenges. Quoting Thomas A. Einhorn, when conducting animal studies, one

has to keep in mind that “in general, the best model system is the one which most

closely mimics the clinical situation for which this technology is being developed,

will not heal spontaneously unless the technology is used, and will not heal when


another technology is used if that technology is less advanced than the one being

tested” [261]. The most effective animal models will therefore (1) provide close

resemblance of the clinical and biological environment and material properties, (2)

encompass highly standardised measurement methods providing objective

parameters (qualitative and quantitative) to investigate the newly formed bone tissue

and (3) are able to detect and predict significant differences between the bone tissue

engineering methods investigated [255]. For clinical modelling and efficacy

prediction of the tissue engineering strategy to be translated into clinical application,

up-scaling to large animal models is therefore inevitable. Thereby, the tissue

engineering therapy can be delivered in the same (or similar) way in which it will be

delivered in clinical settings utilising surgical techniques that match (or closely

resemble) clinical methods at the site that matches the setting in which it will be used

later as closely as possible [255]. The advantage of large animal models (using

nonhuman primates, dogs, cats, sheep, goats, pigs) is the closer resemblance of

microarchitecture, bone physiology and biomechanical properties in humans. They

encompass a well-developed Haversian and trabecular bone remodelling, have

greater skeletal surface to volume areas, show similar skeletal disuse atrophy, enable

the use of implants and techniques similar to the ones used in humans and show

highly localised bone fragility associated with stress shielding by implants. However,

the use of large animal models has disadvantages as well, including the high cost and

maintenance expenses, extensive housing and space requirements, relatively long life

spans and lower bone turnover rates (making longer study periods necessary),

difficulties in standardisation to generate large, homogenous samples for statistical

testing as well as various ethical concerns depending on the species used (e.g.

primates) [259]. Despite several disadvantages, it is inevitable to perform the final

pre-clinical in large animals, as realistically as possible, with relevant loading

conditions and with similar surgical techniques as used in the final procedure in

humans [260]. Large animal models provide mass and volume challenges for

scaffold-based tissue engineering and require surgical fixation techniques that cannot

be tested either in vitro or in small animal models [65]. In general, preclinical

translation testing is performed in large skeletally mature animals, the species most

utilised are dog, sheep, goat and pig [255, 262]. If sufficient preclinical evidence for

the efficacy and safety of the new bone tissue engineering system has been generated

utilising large animal models, clinical trials care undertaken to prove clinical


significance and safety, ultimately leading to the translation of the technology into

routine clinical practice.

2.6.3 Taking composite scaffold based bone tissue engineering from bench to

bedside

In accordance with the above outline rationale for translating bone tissue

engineering research into clinical applications, during the last decade our

interdisciplinary research team has focussed on the bench to bedside translation of a

bone tissue engineering concept based on slowly biodegradable composite scaffolds

made from medical grade polycaprolactone (mPCL) and calcium phosphates

[hydroxyapatite (HA) and tricalcium phosphate (TCP)] [148, 263]. Detailed

descriptions of the scaffold fabrication protocol can be found in recent publications

[75, 176, 263-265].

The scaffolds have been shown in vitro to support cell attachment, migration

and proliferation; degradation behaviour and tissue in-growth has also been

extensively studied [266-269]. We subsequently took the next step towards clinical

translation by performing small animal studies using rat, mice and rabbit models

[270-272]. As reviewed in detail in [263], we were able to demonstrate the in vivo

capability of our composite scaffolds in combination with growth factors or cells to

promote bone regeneration within ectopic sites or critical sized cranial defects in the

small animal models. Studies in large animal models that closely resemble the

clinical characteristics of human disease, with respect to defect size and mechanical

loading, then became essential to advance the translation of this technology into the

most difficult and challenging clinical applications in orthopaedic tumour and trauma

surgery. The choice of a suitable large animal model depends on the ultimate clinical

application, and consequently there is no such things a “one gold standard animal

model”. Over the last years, our research team has investigated the application of our

composite scaffolds in several preclinical large animal models addressing different

clinical applications:


2.6.4 Load-bearing, critical-sized ovine tibial defect model

Well-characterised, reproducible and clinically relevant animal models are

essential to generate proof-of-principle pre-clinical data necessary to advance novel

therapeutic strategies into clinical trial and practical application. Our research group

at the Queensland University of Technology (QUT; Brisbane, Australia) has spent

the last 5 years developing a world-leading defect model to study pre-clinically

different treatment options for cases of large volume segmental bone loss [77, 273].

We have successfully established this 3 cm critical-sized defect model in sheep tibiae

to study the mPCL-TCP scaffold in combination with cells or growth factors

including bone morphogenic proteins (BMPs) [76, 274]. This model has not only

generated a series of highly cited publications [74, 76, 274-276] but also has attracted

large interest in the orthopaedic industry to be used as a preclinical test bed for their

bone graft products under development. The model enables control of experimental

conditions to allow for direct comparison of products against a library of benchmarks

and gold standards we have developed over the last 5 years (we have performed

more than 200 operations using this model to-date). Our preclinical tibial defect

model developed at QUT is one of the only available models internationally, which

is suitable from both reproducibility and cost point of view for the evaluation of large

segmental defect repair technologies in statistically powered study designs. We have

chosen this critical sized segmental defect model of the tibia for our large animal

model because tibial fractures represent the most common long bone fractures in

humans and are often associated with significant loss of bone substance [6, 10]. Also,

tibial fractures result in high rates of non-unions or pseudarthroses [10, 16]. From an

orthopaedic surgeons point of view it can be argued that amongst all bone defects

seen in the clinical practice, segmental defects of the tibia are often the most

challenging graft sites. This owes to the grafts being required to bear loads close to

physiological levels very soon after implantation, this is despite internal fixation,

which often provides the necessary early stability, but also suffers from the poor soft

tissue coverage (vascularisation issue) of the tibia compared to the femur. Hence, in a

bone engineering strategy for the treatment of segmental tibial defects, the scaffold

must bear (or share) substantial loads immediately after implantation. The scaffold’s

mechanical properties (strength, modulus, toughness, and ductility) are determined

both by the material properties of the bulk material and by its structure

(macrostructure, microstructure, and nanostructure). Matching the mechanical


properties of a scaffold to the tibial graft environment is critically important so that

progression of tissue healing is not limited by mechanical failure of the scaffold prior

to successful tissue regeneration. Similarly, because mechanical signals are important

mediators of the differentiation of cell progenitors, a scaffold must create an

appropriate stress environment throughout the site where new tissue is desired.

Hence, one of the greatest challenges in scaffold design for load bearing tibial

defects is the control of the mechanical properties of the scaffold over time. By

trialing our bone tissue engineering strategies in a tibial defect model, we will

therefore addressing a highly relevant clinical problem and are creating valuable pre-

clinical evidence for the translation from bench to bedside. With the 3cm critical

defect being regenerated successfully by applying our mPCL-TCP scaffold in

combination with BMP [75], we are now investigating bone regeneration potentials

in even larger sized tibial defects (Figure 7).

Figure 7: Load-bearing critical-sized ovine tibial defect model using mPCL-TCP scaffolds

manufactured by FDM. Scaffolds (A=clinical image, holes are oriented towards neurovascular

bundle to further promote ingrowth of vasculature) exhibit mechanical and structural properties

comparable to cancellous bone and can be produced with distinct control over scaffold properties

(porosity, pore size, interconnections etc.) by AM. B= Side and top view of a mPCL-TCP scaffold

visualised by microcomputed tomography. The fabrication via FDM enables well-controlled

architecture as evidenced by the narrow filament thickness distribution, leading to a porosity (volume

fraction available for tissue ingrowth) of 60%, with interconnected pores. Scale bars are 5 mm. [Image

B reproduced with permission from (246), © The Authors.] C-H = Surgical procedure: A 6cm tibial

defects is created in the tibial diaphysis (C-D) and the periosteum is removed from the defect site and

additionally also from 1cm of the adjacent bone proximally and distally. Special care is taken not to

damage the adjacent neurovascular bundle (E, bundle indicated by Asterisk). The defect site is then

stabilised using a 12 hole DCP (Synthes) (F). Afterwards 6cm mPCL-TCP scaffold loaded with PRP


and rhBMP-7 is press fitted into the defect site to bridge the defect (G-H) and the plate is fixed in its

final position. Xray analysis at 3 months after implantation (I) shows complete bridging of the defect

site with newly formed radio-opaque mineralised tissue (in order to provide sufficient mechanical

support, the scaffold is not fully degraded yet and scaffold struts appear as void inside the newly

formed bone tissue).

2.6.5 Minimally-invasive ovine thoracic spine fusion model

Spinal fusion has been investigated in animal models for one hundred years

now and a lot of the knowledge we have today on how spinal fusion progresses was

gained through animal models [277, 278]. With regards to the above pictured

rationale for translating bone tissue engineering approaches to clinical practice, it is

of importance to note that the physical size of the sheep spine is adequate to allow

spinal surgery to be carried out using the same implants and surgical approaches that

are used in humans as well. Also, sheep spines allow for an evaluation of the success

of the study using fusion assessments commonly used in clinical practice. When

considering spinal fusion in large animal models, it is apparent that due to the

biomechanical properties of the spine a biped primate animal model (such as in

[279]) should ideally preferred over a quadruped large animal model (for example

ovine[280] or porcine[281]). But given the expenses and limited availability of

primate testing as well as ethical concerns due to the close phylogenical relation, it is

more feasible to trial large numbers of scaffold variations in the most appropriate

quadruped large animal models and then evaluate the best performing scaffold in a

primate model, if possible [65].

We have outlined above that defining specific clinical target applications is a

critical prerequisite for successful bone tissue engineering research that is meant to

be translated into clinical practice. In accordance with this we have selected the

thoracic spine for our animal model because we have identified idiopathic scoliosis

as clinically highly relevant thoracic spine pathology. Idiopathic scoliosis is a

complex three-dimensional deformity affecting 2-3% of the general population

[282]. Scoliotic spine deformities include progressive coronal curvature,

hypokyphosis or lordosis in the thoracic spine and vertebral rotation in the axial

plane with posterior elements turned rotated toward the curve concavity.

Scoliotically deformed vertebral columns are prone to accelerated intervertebral disc


degeneration, initiating more severe morphological changes of the affected vertebral

joints and leading to chronic local, pseudoradicular, and radicular back pain [283].

One of the critical aspects in surgical scoliosis deformity correction is bony fusion to

achieve long-term stability [284]. Autologous bone grafting is still the gold standard

to achieve spinal fusion and superior to other bone grafts for spinal fusion [285-287].

Nonetheless, the use of autologous bone grafting material has significant risks as

outlined in detail above. A number of animal models for the use of tissue-engineered

bone constructs in spinal fusion exists [288] and the use of bone morphogenetic

proteins for spinal fusion has been studied extensively [277, 280, 289, 290].

However, to the best of our knowledge, our ovine thoracic spine fusion model

is the first existing preclinical large animal model on thoracic intervertebral fusion

allowing the assessment of tissue-engineering constructs such as biodegradable

mPCL-CaP scaffolds and recombinant human bone morphogenetic protein-2

(rhBMP2) as a bone graft substitute to promote bony fusion (Figure 8) [291]. We

have been able to show that radiological and histological results at 6-months post-

surgery indicated had comparable grades of fusion and evidenced new bone

formation for the mPCL-CaP scaffolds plus rhBMP-2 and autograft groups. The

scaffold alone group, however, had lower grades of fusion in comparison to the other

two groups. Our results demonstrate the ability of this large animal model to trial

various tissue engineering constructs against the current gold standard autograft

treatment for spinal fusion in the same animal. In the future, we will be able to

compare spinal fusion tissue engineering constructs in order to create statistically

significant evidence for clinical translation of such techniques.


Figure 8: The use of mPCL-CaP scaffolds for spinal fusion. (1) (a) Micro-computed tomography

(m-CT) image of a biodegradable mPCL-TCP scaffold, (b) Representative scanning electron

microscopy image at 100xmagnification. (2) Image of scaffold prior to implantation. (3) Pictorial

series demonstrating the implantation process of a PCL-based scaffold: (a) Cleared intervertebral disc

space prepared for implantation, (b) Implantation process of scaffold into prepared intervertebral disc

space. Scaffold being inserted into prepared intervertebral space, (c) Scaffold in situ within a

predefined intervertebral disc space, (d) Internal fixation with a 5.5mm titanium rod and two vertebral

screws stabilize the treatment level. (4) Representative reconstructed parasagittal CT images at 6

months demonstrating radiologically evident high fusion levels of (a) the recombinant human bone

morphogenetic protein-2 (rhBMP-2) plus calcium phosphate (CaP)-coated PCLbased scaffold and (b)

autograft groups, while lower fusion levels were seen in the (c) CaP-coated PCL-based scaffold alone

group. (5) Representative histological (longitudinal) sections of specimen at 6 months post surgery

from PCL-based scaffold plus rhBMP-2 group exhibiting well aligned columns of mineralized bone

(indicated by letters ‘‘col’’) seen interdigitating with struts of the scaffold filaments (indicated by

letters ‘‘SC’’). Reproduced with permission from (291), © Mary Ann Liebert, Inc.

2.6.6 Current clinical applications of the composite scaffolds and future outlook

The interdisciplinary research group has evaluated and patented the parameters

necessary to process medical grade polycaprolactone (mPCL) and mPCL composite

scaffolds (containing hydroxyapatite or tricalciumphosphate) by fused deposition

modelling [165]. These “first generation scaffolds” have undergone more than 5

years of studies in clinical settings and have gained Federal Drug Administration

(FDA)-approval in 2006 and have also been successfully commercialised

(www.osteoporeinternational.com). The scaffolds have been used highly successfully

as burr whole plugs for cranioplasty [292] and until today more than 200 patients

have received burr whole plugs, scaffolds for orbital floor reconstruction and other

cranioplasties (Figure 9) [160]. With their extensive, multidisciplinary approach the


research team has achieved one of the rare examples of a highly successful bone

tissue engineering approach bridging the gap between scientific research and clinical

practice leading to significant innovations in clinical routines.

As shown above, “second generation scaffolds” produced by FDM and based

on composite materials have already been broadly studied in vitro plus in vivo in

small animal models and are currently under preclinical evaluation in large animal

studies conducted by our research group. Available data so far clearly supports the

view that further translation into clinical use will take place and that a broad

spectrum of targeted clinical applications will exist for these novel techniques.

Our results are consistent with the results of other members of the (bone) tissue

engineering community all around the world, clearly showing the significance of

innovations in the field of tissue engineering. In 2006 Chris Mason proposed two

Figure 9: Clinical case showing the craniofacial scaffold applications for orbital floor fractures.

Moldable scaffolds (A–D) are used and mechanical stability, early vascularisation, osteoconductivity

and ease of handling have been well balanced in the design of mPCL scaffold sheets in order to

properly meet the clinician's needs. The clinical follow up 2.5 years postsurgery (lower CT image) of

a patient receiving a mPCL scaffold (defect site shown in upper CT image) for the reconstruction of a

orbital floor fracture defect showed complete bone regeneration of the defect site (arrow). Reproduced

with permission from (160), © 2007 John Wiley and Sons


distinctly different periods of the regenerative medicine industry, namely,

Regenerative Medicine 1.0 spanning 1985–2002, and Regenerative Medicine 2.0

commencing in approximately 2006 [79]. We herein propose that Regenerative

Medicine 3.0 has commenced. We foresee that the complexity and great variety of

large bone defects require an individualized, patient-specific approach with regards

to surgical reconstruction in general and implant/tissue engineering selection in

specific. We advocate that bone tissue engineering and bioengineering technology

platforms, such as additive manufacturing approaches can be used even more

substantially in bone grafting procedures to advance clinical approaches in general

and for the benefit of individual patient in particular.

The tremendous advantage of scaffolds made by Additive Manufacturing

techniques such as Fused Deposition Modeling (FDM) is the distinct control over the

macroscopic and microscopic shape of the scaffold and thereby control over the

shape of the entire TEC in total. Additive manufacturing enables the fabrication of

highly structured scaffolds to optimise properties highly relevant in bone tissue

engineering (osteoconductivity, osteoinductivity, osteogenicity, vascularisation,

mechanical and chemical properties) on a micro- and nanometre scale. Using high-

resolution medical images of bone pathologies (acquired via CT, µCT, MRI,

ultrasound, 3D digital photogrammy and other techniques) [234] we are not only be

able to fabricate patient-specific instrumentation [293-295], patient-specific

conventional implants [296-300] or allografts [301], but also to realise custom-made

tissue engineering constructs (TEC) tailored specifically to the needs of each

individual patient and the desired clinical application [234, 240, 302]. We therefore

predict that the commencing area of Regenerative Medicine 3.0 will hold a

significant leap forward in terms of Personalised Medicine.

We have already proven the clinical application of this concept by fabricating a

custom-made bioactive mPCL-TCP implant via CAD/FDM that was used clinically

to successfully reconstruct a complex cranial defect [303]. We have also recently

provided a rationale for the use of CAD/FDM and mPCL-TCP scaffolds in

contributing to clinical therapy concepts after resection of musculoskeletal sarcoma

(Figures 10 and 11) [304]. Although it has to be mentioned that our approaches

presented in this review are at different stages of clinical translation, their entity

clearly represents a promising and highly significant 21 century approach in taking


bone tissue engineering strategies from bench to bedside and into the era of

Regenerative Medicine 3.0.

In conclusion, the field of bone tissue engineering has significantly changed the

millennia old quest by humans to optimise the treatment of bone defects and to

identify suitable bone substitute materials. We have reviewed the historic

development, current clinical therapy standards and their limitations as well as

currently available bone substitute materials. We have also outlined current

knowledge on scaffold properties required for bone tissue engineering and the

potential clinical applications as well as the difficulties in bridging the gap between

research and clinical practice. Although the clinical translation of these approaches

has not taken place on a large scale yet, bone tissue engineering clearly holds the

potential to overcome historic limitations and disadvantages associated with the use

of the current gold-standard autologous bone graft. Optimizing combinations of cells,

scaffolds, and locally and systemically active stimuli will remain a complex process

characterized by a highly interdependent set of variables with a large range of

possible variations. Consequently, these developments must also be nurtured and

monitored by a combination of clinical experience, knowledge of basic biological

principles, medical necessity, and commercial practicality. The responsibility for

rational development is shared by the entire orthopaedic community (developers,

vendors, and physicians). The need for objective and systematic assessment and

reporting is made particularly urgent by the recent rapid addition of many new

options for clinical use. By applying a complex interplay of 21st century

technologies from various disciplines of scientific research, the gap between bone

tissue engineering research and the translation into clinically available bone tissue

engineering applications can successfully be bridged.


Figure 10: Clinical case of a 52 year old man with a malignant bone tumour above his left hip.

(A) X ray and computed tomogram showing mixed lytic sclerotic lesion above the left acetabulum,

Technitium-MDP bone scan demonstrating focal area of increased tracer uptake within the tumour.

(B) Tumour resection leaving a large pelvic defect (white arrows), f= femoral head. (C) Resected

specimen including upper part of acetabulum (Clinical images: P.F.C.). The surgical resection creates

a large bone defect in the pelvis that necessitates the use of autograft/allograft bone material and/or

orthopaedic implants to reconstitute the pelvic anatomy. A novel approach (D-G) could be the use of

custom made porous bone tissue engineering scaffolds fabricated via Computer Aided Design (CAD)

to regenerate such defects: Data obtained from high-resolution CT can be used to create a 3D

computer-aided designed (CAD) model of the patient's pelvis by additive manufacturing (D). This

model can be used by the orthopaedic surgeon to indicate osteotomy planes to achieve tumour free

margins, after which, after which the CAD model is virtually resected (E). A custom made scaffold to

fit the defined defect is then created by mirroring the healthy side of the pelvis, adjusting the size of

the scaffold accordingly and fabricating the scaffold from the virtual model using AM techniques (F).

Flanges, intramedullary pegs and other details can be added to the porous scaffold structure to

facilitate surgical fixation and to enhance its primary stability after implantation (G). Images D-G

reproduced with permission from (304), copyright: the authors.


Figure 11: The vascularised fibula transfer combined with bone tissue engineering applications.

The vascularised fibula transfer is one of the most commonly used techniques for reconstruction of

large tibial defects in orthopaedic oncology. The figure shows clinical case from a 16 year old girl

with a malignant tumour of the mid-shaft of tibia. (a) Xray showing destructive lesion. (b) Segmental

resection of tumour. (c) Defect created by the removal of tumour. (d) Example of reconstruction using

vascularised fibular within allograft bone (Cappanna procedure), (e) Reconstruction in-situ using

vascularised fibular and allograft. (f) postoperative X-ray images. (g) 3D computed tomogram of

reconstruction showing fibula enclosed by allograft bone material (Clinical case: P.F.C.). A novel

biological approach to avoid the use of allograft material could be the combination of a vascularised

fibula transfer with a custom made tissue engingeering construct as shown in H: After resection of the

malignant tumour (1), a customized tubular scaffold is placed around the vascularised fibula autograft

to fill the defect (2-3). Primary stability and even load distribution is achieved by using an internal

fixation device (4). Secondary stability is achieved by osseointegration of both the fibula and the

porous tissue engineering scaffold. Over time, the scaffold is slowly replaced by ingrowing tissue

engineered bone and the defect is completely bridged and regenerated (5). H partly reproduced with

permission from (304), © The Authors.

Chapter 3: A novel spatiotemporal delivery system of rhBMP-2 to regenerate critical sized tibial defects in an

ovine large animal model 57

Chapter 3: A novel spatiotemporal delivery 1

system of rhBMP-2 to regenerate 2

critical sized tibial defects in an 3


3.1 INTRODUCTION 5

First described by Hurley et al 1959 for experimental spinal fusion treatment [305], 6

the technique of “Guided Bone Regeneration” (GBR) is currently mainly used in oral 7

and maxillofacial surgery (e.g. in implant dentistry) in humans [306, 307]. The basic 8

principle of GBR is the use of a mechanical barrier membrane (or similar structure) to 9

inhibit or delay migration of cells impeding bone formation (e.g. epithelial cells and 10

fibroblasts) from surrounding tissues into the defect site, thereby favouring migration 11

of pluripotential and osteogenic cells (e.g. osteoblasts derived from the periosteum 12

and/or adjacent bone and/or bone marrow) from adjacent periosteum or bony margins 13

into the defect [306]. However, the use of barrier membranes for guided bone 14

regeneration and restoration of large bone defects has also become a field of interest 15

for various orthopaedic conditions (including revision surgeries and limb salvage 16

procedures) [308]. Although various studies have addressed the use of bioresorbable 17

membranes for reconstruction of segmental mandibular defects in small and large 18

animal models so far, only a small number of studies exists investigating their use for 19

long bone defects in small animal models (see [308] and references therein). And 20

there are only a handful of large animal studies using GBR and resorbable membranes 21

for long bone defect reconstruction: Rhodes and colleagues investigated the use of 22

Hyalonect (membrane comprising knitted fibres of esterified hyaluronan) to cover 23

defects made in the humeri of dogs filled with different bone graft test materials 24

[309]. They found that Hyalonect allowed regeneration of bone within the humeral 25

defects whilst preventing fibrotic tissue in-growth and found regeneration of tissue 26

began to resemble natural periosteal tissue. However, the membranes were only used 27

to cover relatively small circular drill holes (9mm diameter). Oh et al evaluated 28

evaluate the effect of betatricalcium phosphate and poly L-lactide-co-glycolide-29

coepsilon-caprolactone (TCP/PLGC) membrane in the repair of partial bone defects 30

58 Chapter 3: A novel spatiotemporal delivery system of rhBMP-2 to regenerate critical sized tibial defects in an

ovine large animal model

in the canine proximal humerus (length of the defect was quarter of the full bone, and 31

width of the defect was quarter of middle diameter of the lateral aspect of the bone) 32

[310]. It was shown that the TCP/PLGC membrane inhibited fibrous connective tissue 33

migration into the defect site and new cortex growth was present in the defect. The 34

authors concluded that the TCP/ PLGC membrane was a good guided bone 35

regeneration material to restore the original morphology of partial (not full segmental) 36

humerus defect. Beniker et al applied an acellular dermal matrix (GraftJacket 37

Acellular Periosteum Replacement Scaffold, Wright Medical Technology, Inc, 38

Arlington, Tenn) in a porcine midshaft critical-sized femoral segmental defect model 39

[311]. The authors state a defect size “in length two times the diameter of the bone”. 40

The porcine dermal membrane was wrapped around the cylindrical bone defect 41

creating a tube that was was filled with a 1:1 ratio of OsteoSet Pellets (Wright 42

Medical Technology, Ine) mixed with cancellous autograft bone chips obtained from 43

the proximal humerus. New bone formation within the margins of the defect and 44

adjacent to the scaffold was found with minimal to no soft tissue invasion. The study 45

provided preliminary evidence that the dermal membrane material may be used as a 46

scaffold for periosteum regeneration by allowing for cellular repopulation, 47

revascularization, and bone defect restoration. Gerber and Gogolewski used a 48

challenging 7cm tibial diaphysial defect model in sheep (stabilized with 49

intramedullary nailing) to trial the use of resorbable poly-L/DL-lactide membranes 50

(with or without perforations) used empty or combined with autologous bone graft 51

(ABG) or a vascularized periosteal flap [312]. The authors showed that using a 52

perforated membrane combined with ABG as well a perforated membrane covered 53

with a vascularized periosteal flap (and no ABG) led to rapid and stable defect 54

regeneration after 16 weeks. Study groups using an empty perforated membrane 55

without ABG or using a non-perforated membrane with ABG did not heal. A 4cm 56

diaphyseal segmental tibial defect model in sheep was used by Gugula et al in two 57

studies with bioresorbable poly /DL-lactide membranes in a single or double-tube 58

design combined with or without cancellous ABG. In groups without autologous bone 59

graft non-union developed and persisted until the end of the studies. Defect healing 60

was only observed when membranes were combined with cancellous ABG in single- 61

or double-tube technique. The formation of a neo-cortex was shown (with a thickness 62

corresponding to the thickness of the intact cortex) when the cancellous ABG was 63



placed in the space between the internal and external membranes in a double-tube 64

technique [313, 314]. 65

Recently, nanofiber based scaffolds are being explored as scaffolds for TE 66

applications with potential applications in guided bone regeneration as well. The 67

fibrillar structure of collagen is important for cell attachment, proliferation, and 68

differentiation function in tissue cultures and mimicking its structure may lead to 69

engineered tissue more closely resembling native tissues. Polymer nanofibers are an 70

important class of nanomaterials which are focused during the last ten years in the 71

field of TE. Nanostructured materials are smaller in size falling around 1–100 nm 72

range and have specific properties and functions related to the size of the materials. 73

The development of nanofibers has enhanced the scope of fabricating scaffolds to 74

mimic the architecture of natural human tissues at nanoscale. The large surface area to 75

volume ratio of nano- and microfibers combined with its porous structure favours cell 76

adhesion, proliferation, migration, and differentiation; all of which are desired 77

properties for engineering tissues. The high porosity of nanofiber scaffolds provides 78

more structural space for cell accommodation and facilitates efficient exchange of 79

nutrient and metabolic waste between a scaffold and the environment. These features 80

of nanofiber scaffolds are morphologically and chemically similar to the extracellular 81

matrix (ECM) of natural tissue, which is characterized by a wide range of pore 82

diameter distribution, high porosity, effective mechanical properties, and specific 83

biochemical properties [315]. Since nanofiber meshes possess nano-scale features 84

which may provide enhanced cellular response compared to solid walled scaffolds 85

[316], their use as membranes for guided bone regeneration may be advantageous. 86

There are several scaffold fabrication techniques namely, electrospinning, self-87

assembly, phase separation, melt-blown, and template synthesis. Of these techniques, 88

electrospinning is the most widely used technique and also demonstrates most 89

promising results for TE applications [317, 318]. Electrospun nanofiber meshes 90

demonstrate small pore size, flexibility, high porosity, high surface area, and excellent 91

mechanical strengths; therefore, they have been investigated in a wide variety of 92

biomedical applications, including the production of scaffolds for tissue engineering, 93

wound healing, drug delivery, and medical implants [315, 319]. Electrospinning is a 94

relatively simple process and can generate nanofiber meshes with high porosity, large 95

surface area-to-volume ratios, and therefore is an attractive technique for the 96

production of scaffolds [320-322]. The high surface area allows for greater cellular 97



attachment, as well as multiple focal adhesion points on different fibres that enables 98

spreading of the cell in its natural state. Studies have shown that nanofiber meshes are 99

able to support attachment, proliferation and metabolism of numerous cell types 100

including osteoblasts, mesenchymal stem cells, endothelial cells, fibroblasts and 101

primary human haematopoietic stem cells [323] [324] [325-327]. 102

103

Although nano- and microfiber meshes / membranes could themselves be used as 104

functionalized barrier membranes in GBR-applications, studies (including ovine 105

animal studies shown above using autologous bone graft) indicate that for a strong 106

differentiation response an osteoinductive growth factor (either endogenously derived 107

from autologous bone graft material or via externally added growth factors such as 108

BMP) may still be needed. 109

Bone morphogenetic proteins (BMPs) are widely used in Bone Tissue Engineering 110

(BTE) applications as well as in current clinical practice [328-330]. But despite their 111

promising beneficial effects on bone regeneration, several significant issues remain 112

for the routine application of BMPs. Due to suboptimal delivery methods, poor site-113

directed control as well as temporal dosage control, and short protein half-life vastly 114

supraphysiological concentrations of BMPs need to be administered in order for the 115

growth factors to exert their bone regenerative effects in vivo [331-333]. Exogenous 116

BMP is often administered in the range of milligram quantities while localised 117

endogenous BMP production is physiologically at a nanogram range [334]. The 118

current standard dosage of 3.5mg rhBMP-7 used in clinical applications is estimated 119

to be equivalent to twice the amount of the total BMP-7 found in the entire human 120

skeleton [335]. High doses of BMPs have been linked to various side effects 121

including hypertrophic or ectopic bone formation, immunological reactions as well as 122

a potential correlation between extremely high doses of BMP and cancer incidence 123

[336, 337]. Furthermore, high costs associated with large quantities of BMP 124

administered per patient currently limit the routine use of BMPs [333, 338]. Clearly, 125

novel delivery systems providing controlled extended release as well as spatial 126

retention of BMPs are needed to improve efficacy of these growth factors at lower 127

doses and minimise side-effects, increase cost-effectiveness and enable a routine 128

clinical use of BMPs [331, 332]. 129

A variety of naturally occurring and/or synthetic carriers have been investigated by 130

scientific research over the last decades for controlled growth factor delivery 131



including BMPs (as well as for delivery of other proteins) [331, 339, 340]. Most 132

carriers release BMP through an initial high burst, and soon-after the total 133

concentration of BMP drops below therapeutic levels as they lose their biological 134

activity through degradation in the body. This often leads to initial unacceptably high 135

concentrations and poor targeting of the BMP to the tissue of interest, which can lead 136

to severe side effects, low efficiency and non-sustained BMP levels [331, 332]. 137

Alginate, a naturally occurring polymer derived from brown algae such as Laminaria 138

hyperborea, possesses several beneficial characteristics that make this polysaccharide 139

a promising target for the development of delivery vehicles for sustained release 140

applications in tissue engineering (TE) [341]. A critical advantage is the gelling 141

behaviour which allows gentle encapsulation of various substances [342]. The 142

encapsulation and release of proteins, such as BMPs, from alginate gels can 143

significantly enhance their efficacy and targeting. Mammalian cells do not have 144

receptors for alginate polymers, which makes alginate gels themselves relatively inert 145

[341]. However, coupling the fibronectin-derived adhesion peptide arginine-glycine-146

aspartic acid (RGD) and its subtypes to alginate offers a specific way to control cell 147

adhesion as the RGD-sequence is the cell attachment site of a large number of 148

adhesive extracellular matrix, blood, and cell surface proteins and cell receptor-RGD 149

interactions are well characterized [341, 343]. 150

Extensive previous studies from our groups have demonstrated that the in vivo 151

delivery of bone morphogenic protein 2 (BMP-2) within a RGD (arginine-glycine-152

aspartic acid) functionalized alginate hydrogel is a potent technique to stimulate bone 153

formation in a rodent animal model [344-354] (Figures 12+13). 154

Based on the favourable results from the rat femoral segmental defect model, we 155

herein trialled the novel rhBMP-2 hybrid delivery system in our well established and 156

fully characterized QUT 3cm tibial defect ovine animal model [71, 75, 77]. It was 157

hypothesized that rhBMP-2 delivery through the hybrid system in the ovine animal 158

model would lead to significantly increased bone regeneration and improved 159

biomechanical function comparable to our previous results from the rat animal model. 160

By up-scaling from the rodent small animal model to a preclinical ovine large animal 161

model, it was the aim of this study to enable pre-clinical modelling and efficacy 162

prediction of the novel spatiotemporal delivery system for rhBMP-2. 163

164



165

Figure 12: Rat femoral defect model using hybrid delivery system of rhBMP-2. (A) Nanofiber 166 mesh tubes and alginate hydrogel for surgery. SEM image of electrospun nanofiber mesh illustrating 167 the smooth and bead-free nano-scaled fibers. (B) Hollow tubular implant without perforations made 168 from nanofiber meshes. (C) Tubular implant with perforations. (D) Implants in segmental bone defect. 169 Modular fixation plates are used to stabilize the femur. A nanofiber mesh tube is placed around the 8 170 mm defect. In some groups, alginate hydrogel, with or without rhBMP-2 is injected inside the hollow 171 tube. (E) Picture of defect, after placement of a perforated mesh tube. The alginate inside the tube can 172 be seen through the perforations. (F) A specimen was taken down after 1 week and the mesh tube was 173 cut open. The alginate was still present inside the defect, with hematoma present at the bone ends. (G) 174 Alginate release kinetics over 21 days in vitro. Sustained release of the rhBMP-2 was observed during 175 the first week. Reproduced from [349], Copyright © 2010 Elsevier Ltd. All rights reserved. 176 177

178 Figure 13: Representative radiographs at 4 and 12 weeks for rat femoral defect model. Defects in 179 Groups I and II demonstrated small amount of bone formation, and did not bridge, even after 12 weeks. 180 At week 4, defects in Groups III samples were infiltrated with considerable bony tissue, while Group 181 IV samples exhibited the most robust mineralization. All samples in Groups III and IV were bridged 182 with densely packed bone at week 12. Reproduced from [349], Copyright © 2010 Elsevier Ltd. All 183 rights reserved. 184



3.2 MATERIALS AND METHODS 185

3.2.1 Animal Ethics Approval and Code of Practice 186

Study approval was obtained from the Animal Ethics Committee of the Queensland 187

University of Technology (Animal Ethics Approval Number 0900000425). All animal 188

surgeries were performed at the QUT Medical Engineering Research Facility 189

(MERF), The Prince Charles Hospital, Chermside, Brisbane, QLD, Australia. The 190

study was conducted in accordance with all requirements of the Australian Code of 191

Practice for the Care and Use of Animals for Scientific Purposes. 192

193

3.2.2 Scaffold design and fabrication 194

Tubular microfiber scaffolds (Figure 14) were fabricated from medical grade 195

polycaprolactone (mPCL) using direct writing in a melt electrospinning mode as 196

described previously [270, 321, 322]. Scaffold dimensions were set to an outer 197

diameter of 25 mm, a height of 50 mm and an inner diameter of 24 mm. Coating with 198

a calcium phosphate (CaP) layer was added to enhance osteoinductivity of the 199

scaffold in a three step process: Briefly, surface activation with Sodium hydroxide 200

(NaOH) was followed by treatment with simulated Body Fluid 109 (SBF109) to 201

deposit the CaP and post-treatment with NaOH. The coating process has been 202

described in detail elsewhere [214, 322]. Afterwards, scaffolds were stored in a 203

humidity controlled storage chamber until further use. On the day of surgery scaffolds 204

were sterilised by incubation in 70% ethanol for 5 minutes followed by complete 205

evaporation and subsequent UV irradiation for 60 min before implantation. 206

207

3.2.3 Preparation of alginate hydrogel with and without growth factors 208

Gamma irradiated medical grade sodium alginate was covalently coupled with RGD-209

containing G4RGDASSP peptide sequences as previously described [355, 356]. 210

Sodium alginate was then crosslinked at a concentration of 2% (w/v) with calcium 211

sulphate slurry at a ratio of 10:1 to form alginate hydrogels. For treatment group III 212

16.67 μg/ml rhBMP-2 were encapsulated in the alginate hydrogels as previously 213

published [348, 349]. Hydrogels were stored overnight at 4°C prior to injection into 214

implanted tubular scaffolds the following day. 215

216



217

Figure 14: Representative images of tubular microfiber mPCL-scaffolds surface-coated with CaP 218 used in the study. (A) Macroscopic scaffold morphology (B) SEM images in different magnification 219 showing network of smooth electrospun fibers and calcium phosphate surface coating at higher 220 magnifications. Rectangles indicate area of higher magnification shown in following image. 221

222

223

3.2.4 Surgical Procedure 224

In 15 male Merino sheep (50-60 kilogram bodyweight, age ≥6 years) a critical sized 225

3cm tibial defect was created using the surgical technique recently published by our 226

group [75, 77, 273]. In summary a critical-sized 3cm full diameter osteo-peristeoal 227

defect was created in the diaphysis of the right tibia via osteotomy and bone segment 228

removal (Figure 15). Additionally, periosteum was removed circularly over a length 229

of one centimeter from the remaining tibial segments at each osteotomy site. A 230

tubular PCLA-scaffold was then slid 1cm over the end of each osteotomy site of the 231

tibia and fixed in place with purse-string sutures. Afterwards, the defect was 232

stabilized using a 5.6mm 10-hole Dynamic Compression-Plate (DCP, DePuy Synthes) 233

fixed with four 4.5mm cortex screws proximally and three 4.5mm cortex screws 234

distally, respectively. Three different treatment groups (n=5, respectively) were used 235

in this study: In group I the tubular PCLA-scaffold was implanted into the defect site 236

empty (PCLA-only-group). In group II the PCLA-scaffold was combined with 237

injection 6ml functionalized RGD-containing Alginate hydrogel into the scaffold 238

lumen after implantation (PCLA-Alginate-group). In group III (PCLA-Alginate-239



rhBMP-2-group) 6ml of functionalized RGD-containing Alginate hydrogel with an 240

additional 2mg rhBMP-2 were injected into the scaffold lumen after implantation. 241

The wound was then closed in layers. Sterile bandages and a circular cast were 242

applied to the right hind leg. The right hind leg was immobilized in the casts for 4 243

weeks after surgery to reduce load on the operated tibia postoperatively. All animals 244

were kept indoors while casts were present. After cast removal, sheep were released 245

into confined yards and paddocks subsequently. 246

247

248

Figure 15: Surgical procedure. A critical sized 3cm tibial defect was created in the diaphysis of the 249 right hind leg (A-B). A tubular PCL-scaffold was slid 1cm over both ends of the osteotomy sites and 250 fixed with purse-string sutures (C-E). The defect was then stabilized with a 10 hole DC-Plate (DePuy 251 Synthes) (F). In group I, the scaffold was left empty (G). A total of 6ml of functionalized RGD-252 containing hydrogel without (group II) or with (group III) 2mg rhBMP-2 was injected into the scaffold 253 lumen (H-I). Wounds were closed in layers and a circular cast was applied for four weeks to the right 254 hind leg to reduce load on operated tibia. 255

256

3.2.5 Conventional X-ray analysis 257

Immediately after surgery, conventional X-rays in two planes were taken to confirm 258

correct implant (plate, screws) placement and scaffold positioning. Serial 259

conventional X-ray analyses (3.2 mAs, 65kV; Philips, Australia) in two standard 260

planes (anterior–posterior and medial–lateral) were conducted at 3 and 6 months post-261



surgery to assess formation of newly formed mineralised bone tissue and time point of 262

defect bridging. 263

264

3.2.6 Euthanasia of sheep and harvesting of specimens 265

All animals were euthanized humanely at 6 months post-surgery by intravenous 266

injection of 60 mg/kg pentobarbital sodium (Lethabarb; Virbac Animal Health, 267

Milperra, New South Wales, Australia, http://www.virbac.com).) and both hind legs 268

(operated experimental right tibia as well as non-operated control left tibia) were 269

explanted for further analysis in each sheep. Excessive musculature and soft tissue 270

were carefully removed without damaging the defect area. Specimens were frozen at -271

20°C prior to analysis. 272

273

3.2.7 Biomechanical testing 274

DC-plates and screws were carefully removed from experimental tibiae prior to 275

biomechanical testing after resection of bony overgrowth of the plate or screw heads. 276

Bone ends were then embedded in Paladur (Heraeus Kulzer) dental acrylic using 277

custom-made jigs and afterwards mounted in a biaxial testing machine (Instron 8874, 278

Instron, Norwood, USA). Torsion testing was conducted under angular displacement 279

control at an angular velocity of 0.5°/s and a constant compressive preload of 0.05 kN 280

until first signs of fracture occurred. Maximum torsional moment (TM) and torsional 281

stiffness (TS) values were calculated and then normalized against the measured values 282

of the contralateral, non-operated tibia of the same animal. Detailed protocols for 283

biomechanical testing can be found in [75]. After biomechanical testing, all 284

experimental right tibial specimens were cut to a total length of 5cm (complete 3cm 285

defect site length plus 1cm adjacent host bone on each end) and underwent further 286

analysis as listed below. 287

288

3.2.8 Micro computed tomography (micro CT, µCT) 289

After mechanical testing microCT scans of the defect site and adjacent host bone were 290

performed using standardized protocols as recently published by our group [75]. All 291

samples were imaged using a µCT 40 (Scanco Medical AG, Bassersdorf, 292

Switzerland) to quantify newly formed mineralized tissue. Specimens were placed in 293



a sample tube and scanned at an energy of 70 kVp and intensity of 114 µA, resulting 294

in a voxel size of 18 µm. The analysed volume of interest (VOI) included the defect 295

region and adjacent host bone only. Total bone volume (TBV) was measured for the 296

complete defect volume. For comparison of TBV in experimental groups to non-297

operated bones (called “native bone” in the following), six diaphysial tibial bone 298

specimens of 3cm length (explanted from diaphysial parts equivalent to the tibial 299

defect area in the experimental sheep) from male Merino sheep of comparable age 300

and weight not included in this study were scanned using the same parameters and 301

3D-reconstructions as well as TBV were calculated. 302

303

3.2.9 Histology and Immunohistochemistry 304

All analyses were performed in cooperation with the QUT histology laboratory / QUT 305

BTM group and the author would like to acknowledge the great and extensive work 306

of the entire team, especially Flavia Medeiros Savi, Felicity Lawrence and A/Prof 307

Mia Woodruff. After biomechanical testing and microCT analyses, the tibial samples 308

were cut to 5 cm length (3cm defect length plus 1cm of proximal and distal host 309

bone). Afterwards samples were fixed in 10% neutral buffered formalin for 1 week. 310

For histological analysis, samples were then sectioned in transverse and sagittal 311

planes (for cutting schematics please refer to histology figures below). For paraffin 312

format the transverse planes were cut into three sections (P1, P2 and P3). The bone 313

samples were decalcified in 10% EDTA for 6-8 weeks at 37°C using a rapid 314

decalcifier at an input voltage 230V-59Hz, 8A and 450rpm (Kos Milestone 315

microwave model 67051, ABACUS, Brisbane, Australia). The samples were then 316

serially dehydrated in ethanol in a tissue processor (Excelsior ES, Thermo Scientific, 317

Franklin, MA, USA), and embedded in in molten paraffin wax at 60C (Thermo 318

Shandon Histocentre 3 Embedding Station, Thermo Scientific, Brisbane, Australia). 319

10 sections were cut at 5 µm with a Leica RM2235 rotary microtome (Leica 320

Biosystems, Nussloch Germany). Paraffin ribbons were flattened on a water bath 321

(Labec, Marrickville, Australia) at 40C and collected onto polysine microscope slides 322

(Thermo Scientific, Brisbane, Australia) prior to drying at 60C for 16 h. Two slides 323

were then stained with Hematoxylin and Eosin staining (HD scientific, Wetherill 324

Park, Australia) & Eosin (HD scientific, Wetherill Park, Australia) using a Leica 325

Autostainer XL (Leica Biosystems, Nussloch, Germany). The slides were scanned 326



using a Leica SCN 400 slide scanner (Leica Microsystems, Wetzlar Germany) with a 327

20x objective. For resin format samples were sectioned in two sagittal planes (R1 and 328

R2) at 2mm thick slices using an EXAKT 310 Diamond Band Saw (EXAKT 329

Apparatebau GmbH & Co.KG, Norderstedt, Germany). Following degreasing with 330

xylene, the allocated samples for Technovit 9100 New® were processed and 331

embedded in the low-temperature embedding system Technovit 9100 New® (Heraeus 332

Kulzer GmbH, Germany). For ground section format the mounted resin blocks were 333

sectioned longitudinally at 200µm using a EXAKT 310 Diamond Band Saw and 334

subsequently ground at 50µm using a EXAKT 400CS micro grinder (EXAKT 335

Apparatebau GmbH & Co.KG, Norderstedt, Germany) according to the technique 336

described in Donath 1995. Histological assessment was performed using Goldner’s 337

trichrome staining. For thin sections format samples were sectioned with sledge 338

microtome (Polycut-S, Reichert-Jung, International Medical Equipment, USA) using 339

a tungsten carbide blade at 6 μm. Sections were then flattened with 95% ethanol onto 340

Gelatin-coated microscope slides. Following stretching, sections where then covered 341

with polyethylene film and compressed on a benchtop paper to remove ethanol 342

excess. Slides sections were stacked in a metal slides holder to dry for 3-4 days at 60 343

°C. Samples were then stained with von Kossa/McNeal’s Tetrachrome to identify new 344

bone formation. For immunohistochemistry, paraffin sections were deparaffinised 345

with xylene and rehydrated with serial concentrations of ethanol. Subsequently, 346

sections were rinsed in distilled water and placed in 0.2 M Tris-HCl buffer (pH 7.4). 347

Endogenous peroxidase activity was blocked by incubating the sections in 3% H2O2 348

in Tris-HCl for 20 min. This was followed by three washes with Tris buffer (pH 7.4) 349

for 2 min each. Sections were incubated with Proteinase K (DAKO, Botany, 350

Australia) for 20 min and subsequently incubated with 2% bovine serum albumin 351

(BSA) (Sigma, Sydney, Australia) in DAKO antibody diluent (DAKO) in a 352

humidified chamber at room temperature for 60 min to block non-specific binding 353

sites. Afterwards, immunohistochemical staining was performed using primary 354

antibodies specific to the osteogenic markers: 355

356

1. Type I collagen Ab 34710 dilution 1:100 Dab: 2 min (rabbit polyclonal 357

Abcam, Cambridge, UK). 358

2. Bone Morphogenetic Protein 2&4 SC 137087 dilution 1:50 Dab: 1:30 min 359

(Santa Cruz, Biotechnology, CA, USA). 360



3. Von Willebrand Factor A0082 dilution 1:700 Dab: 1:30min (Ready to use, 361

rabbit polyclonal, IR527, Dako, Glostrup, Denmark) 362

4. Cluster of differentiation 31(CD31) SC 1506R dilution 1:1000 Dab: 1:30min 363

(Santa Cruz, Biotechnology, CA, USA). 364

5. Cluster of differentiation 68 (CD68) Ab 125212 dilution 1:300 Dab: 2min 365

(Abcam, Cambridge, UK). 366

6. Alkaline phosphatase (ALP) Ab 108337 dilution 1:500 Dab: 5min (Abcam, 367

Cambridge, UK). 368

7. Vascular Endothelial growth factor (VEGF) SC 152 dilution 1:500 Dab: 369

1:30min (Santa Cruz, Biotechnology, CA, USA). 370

371

The sections were incubated with the specific antibody in humidified chambers at 4°C 372

overnight. Sections were then washed three times for 2 min with Tris buffer (pH 7.4) 373

and incubated with peroxidase labelled dextran polymer conjugated to goat anti-374

mouse and anti-rabbit immunoglobulins (DAKO EnVision+ Dual Link System 375

Peroxidase, DAKO) at room temperature in humidified chambers for 60 min. Colour 376

was developed using a liquid 3,3-diaminobenzidine (DAB) based system (DAKO). 377

Kaiser’s glycerol gelatin (DAKO) was used for coverslip mounting. 378

379

3.2.10 Statistical analysis 380

Statistical analysis was performed using one-way ANOVA and Tukey’s multiple 381

comparison test (GraphPad Prism 7.02, GraphPad Software Inc.). Differences 382

between groups were considered to be statistically significant at p values <0.05. 383

384



3.3 RESULTS 385

3.3.1 Surgical procedure and postoperative follow-up 386

All 15 sheep tolerated the surgical procedure well. Where applicable, all scaffolds 387

contained the injected 6ml alginate hydrogel well with only minimal leakage. 388

Postoperative recovery was achieved without any complications in all sheep. 389

Postoperative follow-up was without any adverse events. No postoperative infections 390

or other complications were observed. All animals included in this study were in good 391

health and survived the experimental period, gaining weight in the months following 392

surgery. 393

394

3.3.2 Radiographic analysis 395

Correct positioning of the defect site as well as the implants (plate and screws) was 396

confirmed radiographically immediately after surgery in two standard planes 397

(anteroposterior and mediolateral). 398

Two months post-surgery, robust new bone formation was observed for group III 399

(Scaffold+Alginate+rhBMP-2-group) with some animals showing radiographic signs 400

of full bony bridging of the defect site at this early time point. In contrast to this, little 401

or no significant bone formation was observed in group I (Scaffold only-group) and 402

group II (Scaffold+Alginate-group), respectively. No bony bridging of the defect site 403

had occurred in these groups after 2 months. 404

At three month post-surgery (Figure 16, left column) all animals in group III showed 405

radiographic signs of full bony bridging in defect site. Animals in both group I and 406

group II showed sign of beginning bone regeneration mainly originating from the 407

dorsal and proximal part of the tibia (where the defect is covered by the muscles of 408

the lower leg). However, no defect bridging or substantial bone volumes in the defect 409

site could be observed. 410

At the six months-time point (Figure 16, right column) conventional X-ray analysis in 411

two planes revealed complete filling of the defect volume with radio-opaque new 412

bone in all sheep of group III. Early stages of bone-remodelling into cortex and 413

medullary cavity could also be observed originating for the former osteotomy sites. 414

Partial bony bridging was observed in the majority of sheep in group II and to some 415

extent in group I. However, bone volumes in the defect site seemed to be significantly 416



less radiographically and no full bony regeneration of the defect volume was 417

observed. 418

419 Figure 16: Representative clinical radiographic images at 3 and 6 months after surgery. Defects 420 reconstructed with tubular mPCL-scaffold only (Group I), mPCL-scaffold and RGD-Alginate (Group 421 II) and mPCL-scaffold with RDG-Alginate and 1mg rhBMP-2. Complete bony bridging is observed in 422 group III as early as 3 months after surgery, while other groups show almost no radiopaque bone 423 formation at this stage. Complete filling of former defect site due to substantial new bone formation is 424 observed in group III over the course of the study, while the other groups show delayed and only minor 425 bone regeneration without complete bony filling of the defect volume at the end point (6 months). 426 White arrowheads indicate (former) defect size. 427

428



3.3.3 Biomechanical analysis 429

After euthanasia of all animals at 6 month post-surgery left and right tibiae were 430

explanted and biomechanical testing was performed on all specimens. Results for the 431

operated right tibia were compared to the results of the corresponding contralateral, 432

non-operated left tibia for each animal individually. 433

Biomechanical testing showed significantly higher results for maximum torsional 434

moment (TM) and torsional stiffness (TS) in group III (Scaffold+Alginate+rhBMP-2-435

group) compared to group II (Scaffold+Alginate-group) (p = 0.0103 and p = 0.005 for 436

TM and TS, respectively) and group I (Scaffold only-group) (p = 0.0301 and p = 437

0.0192 for TM and TS, respectively) (Figure 17). No statistically significant 438

differences between the group II and group I could be detected for TM and TS 439

(p>0.99 for TM and TS, respectively). 440

441

442

Figure 17: Results of biomechanical testing at 6 months after surgery. No significant differences 443 between group I (scaffold only) and group II (scaffold+alginate) were found. However, TS and TS was 444 significantly higher in group III (scaffold+alginate+BMP2) compared to all other groups. Asterisks 445 indicate statistical significance (p<0.05). 446



3.3.4 Micro computed tomography (µCT) 447

After biomechanical testing all samples were scanned using micro computed 448

tomography (µCT). Three-dimensional (3D)reconstructions of calcified tissue in the 449

defect volume and adjacent host bone confirmed full bridging of the defect site in all 450

samples of the group III (Scaffold+Alginate+rhBMP-2) with substantial new bone 451

formation filling almost the complete defect volume (Figure 18, A). 3D 452

reconstructions for group II and group I (Scaffold+Alginate and Scaffold only, 453

respectively) showed less bone formation than in group III. While bony bridging over 454

the defect length had occurred in most samples of group I and II, the volume of 455

calcified tissue in the defect site seemed to be lower and the tissue appeared to be less 456

remodeled than in group III (Figure 18, B and C). A trend towards higher bone 457

volumes in group II compared to group I was also observed. 458

459

460

Figure 18: Three-dimensional reconstructions of microcomputed tomography (µCT)-scans. 461 Representative three-dimensional reconstructions of micro-CT scans (proximal bone end facing 462 upward) for study groups I – III. Fracture lines visible resulted from biomechanical testing (torsion 463 until failure) before micro-CT analysis. 464

465

Visual results from the 3D reconstructions of the micro-CT scans were confirmed 466

when analyzing Total Bone Volume (TBV = bone volume over the complete defect 467

size) statistically (Figure 19). Mean values of TBV were significantly higher in group 468

A B C



III compared to values in group II (p=0.0092) and group I (p=0.048). Furthermore, 469

statistically significant differences were also found for TBV-comparison between 470

group III and native bone (p=0.023). No statistically significant differences were 471

found between groups I and II and native bone, respectively. 472

473

474

Figure 19: Total Bone Volumes (TBV) (=bone volume over the complete defect size) showed 475 significantly higher TBV for group III (Scaffold+Alginate+BMP2) compared to group I (scaffold) and 476 group II (scaffold+alginate). Furthermore, significantly higher TBV were found for group III in 477 comparison to native bone samples. No statistically significant differences between group I, group II 478 and native bone were found. Asterisk indicates statistical significance (p < 0.05). 479

480

3.3.5 Histology and Immunohistochemistry 481

In accordance with the imaging results obtained from conventional X-rays analyses 482

and microcomputed tomography (mCT), histological analysis confirmed complete 483

defect bridging and formation of mature bone tissue inside the defect volume for all 484

samples in group III as shown by Haematoxylin-Eosin Stain, Goldner’s trichrome 485

Stain and von Kossa/McNeal’s Tetrachrome Stain (Figures 20-22 provide an 486

overview over the histological and immunohistochemical analyses performed for each 487

group). 488

Chapter 3: A novel spatiotemporal delivery system of rhBMP-2 to regenerate critical sized tibial defects in an ovine large animal model 75

489

Figure 20: Overview of results from histological stains and immunohistochemical analysis of group III (Scaffold+alginate+rhBMP-2-group, representative sample specimen). Top left 490 row shows X-ray images at study end point 6 months with schematic of sample explantation and processing. Sagittal plane 3D-reconstruction of microcomputed tomography of corresponding 491 sample shows amount of mineralized tissue with defect margins. Images from undecalcified resin-embedded sagittal sections (as indicated in schematic) stained with Goldner’s trichrome and 492 Kossa/McNeal’s Tetrachrome are shown on top right Bottom row left shows schematic of horizontal sample cuts for decalcification and Paraffin embedding. Representative images for all three 493 defect regions stained with Haematoxylin Eosin (H&E) as well as immunohistochemical analyses (antibody against epitopes listed on top of each column) are shown in the bottom row. 494

76Chapter 3: A novel spatiotemporal delivery system of rhBMP-2 to regenerate critical sized tibial defects in an ovine large animal model

495

Figure 21: Overview of results from histological stains and immunohistochemical analysis of group II (Scaffold+alginate-group, representative sample specimen). Top left row shows 496 X-ray images at study end point 6 months with schematic of sample explantation and processing. Sagittal plane 3D-reconstruction of microcomputed tomography of corresponding sample 497 shows amount of mineralized tissue with defect margins. Images from undecalcified resin-embedded sagittal sections (as indicated in schematic) stained with Goldner’s trichrome and 498 Kossa/McNeal’s Tetrachrome are shown on top right Bottom row left shows schematic of horizontal sample cuts for decalcification and Paraffin embedding. Representative images for all three 499 defect regions stained with Haematoxylin Eosin (H&E) as well as immunohistochemical analyses (antibody against epitopes listed on top of each column) are shown in the bottom row. 500


501

Figure 22: Overview of results from histological stains and immunohistochemical analysis of group I (Scaffold only-group, representative sample specimen). Top left row shows X-ray 502 images at study end point 6 months with schematic of sample explantation and processing. Sagittal plane 3D-reconstruction of microcomputed tomography of corresponding sample shows 503 amount of mineralized tissue with defect margins. Images from undecalcified resin-embedded sagittal sections (as indicated in schematic) stained with Goldner’s trichrome and Kossa/McNeal’s 504 Tetrachrome are shown on top right Bottom row left shows schematic of horizontal sample cuts for decalcification and Paraffin embedding. Representative images for all three defect regions 505 stained with Haematoxylin Eosin (H&E) as well as immunohistochemical analyses (antibody against epitopes listed on top of each column) are shown in the bottom row. 506



Although defect area was completely filled with newly formed bone in group III,

there seemed to be some extent of ongoing remodelling into cortex and medullary

canal extending from both the proximal and the distal osteotomy sites. The mode of

bone formation was found to be mainly direct ossification with osteoid and bone

lining cells present. Newly formed bone was found to be mainly cancellous woven

bone in the centre of the defect with partial remodelling into lamellar bone and

Osteones. Remnants of the injected alginate (Alginate stained blue-purple in H.E.,

red in Von Kossa/McNeal’s Tetrachrome Stain, yellow in Goldner’s trichrome Stain)

as well as bone marrow present in between bone. No alginate was visualized outside

the scaffold dimensions, indicating that the injected alginates most likely had been

contained well by the porous scaffolds (although this was not further analysed

histologically) (Figure 23, A). Compact lamellar bone with primary osteons

including Haversian canals and surrounding interstitial matrix was found to be

present in the peripheral regions of the former cortex, indicating mature and regular

bone formation with potential remodelling towards restoration of a bone cortex.

Medical grade PCL-Scaffold struts (circular voids due to the scaffold dissolving

during tissue preparation in the staining process) were still present after 6 months in

all specimens. Scaffold struts seemed to be fully incorporated into newly formed

bone (including inside osteons) in the neo-cortex (Figure 23, B-C). Further towards

the periphery scaffold struts were surrounded by highly vascularized soft tissue with

an outermost neo-periosteum-layer covering the scaffolds. Osteoid was being

deposited around and inside alginate islets (Figure 23, D). Polynucleated phagozytic

cells were found adjacent to scaffold struts as well as the alginate most likely

degrading and resorbing the hydrogel. Bone marrow including hematopoietic cell

lines and fat cells was present, further indicating mature bone formation and

restoration of bone as a functional organ.


Figure 23: Details of Haematoxylin-Eosin-Stain of representative samples of group III (scaffold + alginate + rhBMP-2). (A) Remnants of alginate (Alg) were found to be closely

surrounded by newly formed mature bone and located inside the scaffold volume (no leakage). Bone marrow (Bm) including fat cells was present indicating regeneration of bone as a fully

functioning organ. Scaffolds struts (Sc) were fully incorporated into new bone or surrounded by highly vascularized soft tissue adjacent to the bone. An outermost neo-periosteum layer (P) was

found to cover the scaffolds. (B) Detailed image showing scaffold (void in image, area of scaffold wall labelled with Sc) incorporation into bone and highly vascularized soft tissue layer (Bv=

Large blood vessel, yellow arrow heads indicate smaller blood vessels) . Bm = Bone marrow. (C) Mature lamellar bone with osteons (Os) including Harversian canals with central blood vessel

(Bv) were found to be present. Furthermore, large blood vessels were present in the neo-periosteal region. (D) Osteoid was found to be deposited around and inside alginate (Alg). Arrows

indicate two osteoblasts invading alginate and depositing osteoid. Bone marrow (Bm) with different cell lines was present indicating mature bone formation.


In group II (Scaffold + Alginate) (Figure 21) partial bony bridging of the defect site

was found to be present. Mature bone as indicated by lamellar structure and

formation of osteons with central Harversian canals was present in parts of the defect

volume as well as bone trabecules. Small islets of enchondral bone formation were

found in advancing hard callus regions (Safranin O/Fast Green Staining). However,

there were also areas inside the scaffold volume where alginate was not incorporated

in bone but surrounded by fibrous vascularized soft tissue. Where new bone had

formed, the scaffold struts were partly incorporated into bone or in more peripheral

regions were embedded in fibrous soft tissue with a high density of surrounding

vasculature. Furthermore, a neo-periosteum-like tissue was found on the outermost

layer covering the scaffolds. In areas where no new bone had been formed, scaffolds

struts were again embedded in fibrous tissue with intersecting blood vessels present

covered by a neo-peristeoum-type of tissue.

In group I (Scaffold only) (Figure 22) smaller amounts of bone volume were present

with partial bridging in some cases, but the main defect volume was found to be

filled with invading soft tissue and muscle (due to collapse of the tubular scaffold

mesh structure over time and loss of barrier function). Ossification was mainly

intramembranous as seen in the other groups. However, areas of enchondral

ossification were present (mainly at the tip of the bone formation advancing from

each osteotomy site into the defect) as shown by Safranin O/Fast Green Stain. Again,

fibrous soft tissue with different size blood vessels formed around the scaffold struts

inside the former scaffold wall.

For further analyses, immunohistochemical staining using primary antibodies against

Type I-Collagen (Col I), Bone Morphogenetic Protein 2&4 (BMP-2/4), Von

Willebrand Factor (vWF), Cluster of differentiation 31 (CD31), Cluster of

differentiation 68 (CD68), Alkaline phosphatase (ALP) and Vascular Endothelial

growth factor (VEGF) was performed:

Newly formed bone in all samples stained strongly positive for Col I as early

osteogenic marker (Figure 24). Collagen Type I was also found to be present in the

fibrous soft tissue adjacent to newly formed bone tissue where the defect site was not

entirely regenerated with bone (mainly in specimens from groups I and II).

Furthermore, fibrous tissue around remaining alginate stained positive for Col I.



Figure 24: Representative sample (group III) of anti-Collagen 1-antibody IHC. Newly formed

bone and interstitial matrix (arrows) as well as soft tissue surrounding alginate (arrowheads) stained

strongly for Col 1. Black bars indicate 100µm.

ALP as a marker for ongoing matrix mineralization was shown to be expressed

strongly at the interface between bone and adjacent soft tissue in the scaffold wall at

the periphery of the samples in groups I, II and III (Figure 25, A-B). Furthermore, in

the centre of the defect area ALP was found in the outer lining of bone trabeculae

were osteoid had been deposited. ALP was also present around the CaP-coated

scaffold struts. In group II and III samples also showed ALP presence at the interface

of bone and soft tissue in the advancing hard callus (Figure 25, C).


Figure 25: Representative samples of IHC using antibody against ALP. (A-B) Sample from group

III showing ALP-presence at interface between bone and soft tissue in scaffold wall (arrow) as well as

in bone trabeculae centrally, where osteoid was being deposited (arrowheads). (C) Image from group I

showing advancing hard callus with positive staining for ALP at bone/soft tissue interface (arrows).

Isotype controls not shown. Black bars indicate 100µm.

As outline above, formation of a highly vascularized fibrous soft tissue between

scaffold struts that were not embedded on newly formed bone tissue had been

observed in histological stains. In accordance with this observation, the

corresponding areas showed profound staining for VEGF, CD 31 and vWF (Figure

24) indicating the presence of vascular growth factors, angiogenesis and endothelial

cells. Positive staining was also found in bone marrow (including larger volume

vessels) and newly formed bone.



Figure 26: Representative images (group III) for IHC using antibodies against VEGF, CD31

and vWF. Scaffold area adjacent to the newly formed bone was found to stain strongly for VEGF,

CD 31 and vWF (arrows), indicating angiogenic signalling as well as presence of endothelial cells

(vasculature). Sc= Scaffold wall area. Black bars indicate 100µm. Isotype controls not shown.

Were scaffold struts were not fully embedded in bone tissue, adjacent soft tissue

areas stained strongly positive for CD 68 in all groups, indicating presence of cells of

the macrophage lineage (most likely showing foreign body reaction to scaffold

surface/material) (Figure 27). CD 68 positive staining was also found in some

samples around scaffold struts embedded in newly formed bone, though staining for

CD68 was generally strongest in soft tissue around scaffold struts. Furthermore, bone

marrow stained also positive for CD68, indicating presence of macrophage lineage

cells including osteoclasts (see overview in Figures 20-22).


Figure 27: Representative images of IHC with anti-CD68 antibody. Scaffold struts appear as

circular void due to dissolving during tissue preparation. (A) Strong CD68-positive staining was

observed in soft tissue directly adjacent to scaffold struts at the interface between bone and soft tissue

(arrows) in the periphery of newly formed bone in all groups (image from group II). (B) Some

samples also showed CD68-positive staining around scaffold struts (arrowheads) embedded in bone

(image from group III), although this was inconsistent and not group-specific. (C) Overview image of

strongly CD68-positive staining around scaffold struts (found along complete length of the scaffold,

indicated by arrows) in all groups (image from group I). Isotype controls were negative for

corresponding areas (data not shown). Black bars indicate 100µm.

In accordance with CD68-results, Tartrate-resistant acid phosphatase (TRAP)-

staining revealed dense presence of TRAP-positive cells in direct contact with

scaffold struts surrounded by fibrous soft tissue in the scaffold wall area (Figure 28,

left column) in all experimental groups. Being CD68-positive, showing TRAP-

staining and being in direct contact with the scaffold material outside the bone tissue

these cells are likely to be activated macrophages interacting with (and resorbing) the

scaffold material. CD68-positive and TRAP-positive osteoclasts were also found

along the newly formed bone tissue in typical locations, indicating presence of a

mature bone with remodelling processes (Figure 28, right column). No obvious

differences in osteoclast location or numbers were found comparing samples form

the three study groups (although no quantitative analysis performed).



Figure 28: Representative images of Tartrate-resistant acid phosphatase (TRAP)-staining. In

accordance with CD68-IHC results, a dens accumulation of TRAP-positive cells was found in direct

contact with the scaffold struts in the soft tissue around newly formed bone (most likely activated

macrophages interacting with scaffold material) (arrows, left column). Furthermore, TRAP-positive

osteoclasts (arrowheads) were found adjacent to newly formed bone tissue in typical locations,

indicating mature bone with ongoing remodelling processes. Black bars indicated 100µm.

Immunohistochemistry using an anti- Bone Morphogenic Protein 2&4-antibody

showed that BMP-2-epitopes were still present in the remaining alginate of group III

six months after implantation (although no conclusion regarding biological activity

could be drawn from this) (Figure 29, A). However, positive staining for BMP-2&4

was also present in bone lining cells as well as their close proximity (Figure 29, B) as

well as around scaffold struts in soft tissue (Figure 29, C) and embedded in bone

(Figure 29, D). These finding were consistent in all groups including group I and II

where no rhBMP-2 had been added exogenously indicating endogenous

osteoinductive BMP2&4-signalling to be present at these sites.


Figure 29: Representative images from IHC using antibody against BMP-2&4. (A) BMP2

epitopes (arrowheads) were found to be still present in remnants of alginate in group III after 6 months

(although no conclusion regarding biological activity could be drawn from this). BMP-2&4-positive

staining was also found in and around bone lining cells (B, arrowheads) as well as around scaffold

struts in soft tissue (C, arrowheads) and embedded in newly formed bone (D, arrowheads). Isotype

controls negative for corresponding areas, images not shown. Black bars indicate 100µm.



3.4 DISCUSSION AND CONCLUSION

Despite multiple innovations over the last decades large bone defects still represent a

major challenge in today’s clinical practice. Large bone defects are associated with a

high rate of pseudarthrotic non-unions, regularly require multiple surgical

interventions (often by several disciplines such as orthopaedic surgery, plastic and

reconstructive surgery, etc.) and often have an unfavourable clinical outcome failing

to restore full function to an injured limb [357]. Therefore, the development of novel

bone tissue engineering applications for the treatment of large segmental bone

defects (especially in load bearing bones) have received considerable scientific,

economic and clinical interest over the past years [61, 62, 358].

In this study, we investigated the regenerative potential of a spatiotemporal delivery

system for reduced doses of rhBMP-2 in our well-established preclinical ovine

animal model with a critical sized 3cm tibial defect. The hybrid delivery system has

previously been successfully applied and extensively characterized by our research

groups using a rat femoral defect model [344-354]. With the current study, we were

able to transfer our previous work utilizing rodent animal models to the application

of this novel tissue engineering strategy in an ovine animal model, taking another

significant step towards clinical translation. We have successfully adapted scaffold

design parameters and upscaled the meltelectrospinning manufacturing process for

the scaffolds to be applicable in our preclinical ovine large animal model. We have

also proven that the total volume of applied alginate hydrogels can be significantly

increased (from µl-range in rodent animal models to a total of 6ml used per sheep

here) with increased size scaffolds still retaining and spatially confining the

functionalized hydrogels well. Furthermore, the tissue engineering constructs (TEC)

have now been implanted using surgical techniques and osteosynthesis material

reflecting current clinical methods in humans. Using an orthotopic sheep model (=

implantation at a skeletal site that matches the setting in which the TEC is supposed

to be used clinically later, i.e. the tibial bone) we were able to assess the regenerative

potential under close resemblance of the situation in human patients (anatomy, bone

physiology and biomechanical properties).

The results of the current study parallel the results obtained from small animal

models in previous studies by our research group and affiliated institutes. Our group


has extensively investigated the novel hybrid alginate hydrogel system in a

preclinical critical-sized rat femoral segmental defect model [344-354]. We

previously found that the application of the hybrid delivery system led to consistent

bony bridging of the segmental defects in the rat model [346, 348, 349, 351, 353].

However, it was also noted that in the absence of rhBMP-2-delivery, a combination

of nanofiber mesh scaffold and alginate alone did not have the capability to

regenerate the critical sized bone defects [348, 349, 353]. We were herein now able

to show the same results for the application of the hybrid system in our ovine large

animal model (Figure 30): The implantation of tubular mPCL-scaffolds only (group

I) or tubular mPCL-scaffold injected with 6ml hydrogel without rhBMP-2 led to

some new bone formation in the defect site. But only the simultaneous delivery of

rhBMP-2 through the hybrid system led to substantial new bone formation with

consistent defect bridging and restoring mechanical properties of the tibia. It seems

that the exogenous addition of growth factors is necessary to bridge such challenging

bone defects in vivo, which is in accordance with one of our previous study [359] and

with recommendations given for current clinical practice even with using additional

autograft bone [360].


Figure 30: Direct comparison of results from rat femoral defect model (left column) and ovine animal model (right column). Results from the preclinical large ovine animal model

directly parallel results from rodent animal model. This is one of the rare cases were a tissue engineering application optimized in a small animal model is apparently as efficient when directly

applied in a preclinical large animal model as well, bridging the scale-up gap between small animal studies and potential clinical translation.


In the present study total bone volumes (TBV) in the defect site of the rhBMP-2-

group (group III) not only were significantly higher compared to the other study

groups (group I and II), but were also found to be significantly higher than

corresponding native tibial diaphysial bone of the same volume. The fact that TBV

were higher than native tibial bone may be explained by the fact that entire defect

was filled with porous bone in the rhBMP-2-group including the former medullary

cavity. While native bone possesses a distinct formation of dense cortex structures

and low-density trabecular bone / a medullary cavity, such remodelling had not taken

place entirely in group III, thereby potentially causing higher total bone volumes to

be present. The total bone volumes in group III of this study are in the range of

results observed in our previous study using a tissue engineering construct with a

FDM-manufactured mPCL-Tricalcium-Phosphate-Scaffold combined with 3,5mg of

rhBMP-7 / PRP after one year [75, 361]. The ABG-control group used in that study

had overall lower TBV than group III from this study, an observation that is

consistent with previous results from the rat femoral defect model [351]. However,

both studies had a 12 month endpoint (whereas a 6 months endpoint was investigated

in this thesis), thereby compromising statistical comparability. Interestingly,

significantly increased total bone volumes in group II compared to group I were also

found, indicating that the presence of the scaffold and functionalized alginate as

matrix does have a positive effect on bone formation in the defect site.

Analysing biomechanical properties we found significantly higher values for

torsional stiffness (TS) and maximum torque (TM) (ultimate torsional strength) in

the rhBMP-2 group (group III) compared to groups I and II. Results were an average

80% TS and 50% TM of contralateral unoperated tibia of the same animal in group

III. These results confirm that the addition of rhBMP-2 in group III not only led to

significantly higher total bone volume values, but also caused a significantly better

restoration of biomechanical properties investigated than the other two study groups.

This finding is also in accordance with previous results from the rat model, where

groups containing rhBMP-2 reached significantly higher biomechanical properties

compared to other study groups [349]. Similar to µCT-results, biomechanical

properties in group III were found to be in the range of results of 3cm-defects treated

with ABG or mPCL-TCP scaffolds and rhBMP-7 after 1 year [75, 361]. Despite a

trend towards higher bone volumes in group II compared to group I, we did not find



statistically significant differences between these groups for the biomechanical

properties analysed. Although the presence of the functionalized alginate hydrogel

led to higher total bone volumes, this bone seemed to be not yet sufficiently

organized to also increase biomechanical properties in group II compared to group I.

With combined histological and immunohistochemical analyses we were able to

proof substantial new bone formation with mineralized matrix in the rhBMP2-group

(group III) compared to the other study groups. Newly formed bone was well

integrated with host bone and highly structured: A cancellous-type woven bone with

trabecular organization was seen in the central defect areas (with some remodelling

into lamellar bone and sparse presence of Osteons as well). Towards the periphery of

the defect, bone was mainly compact lamellar bone with Harversian systems,

indicating formation of a neo-cortex. Reconstitution of bone marrow spaces and

remnants of alginate (where injected) were found similar to observations in the rat

animal model [349]. Furthermore, we found that the periosteum had been

regenerated extending from the host bone towards the defect middle (further data not

shown) forming a neo-periosteal outermost layer. Formation of a periosteal-like

tissue inside the barrier membrane has also been described in another large animal

model with canine humeral defects [309, 311].

Results indicate a complete regeneration of functional bone tissue with periosteum,

potential ongoing reorganization into cortical and cancellous bone as well as

medullary cavity and bone marrow formation / presence (partial macroscopic

remodelling of the defect site into cortex and medullary canal also visible in Figure

20, top row staining). We also found extensive presence of CD68 and TRAP-positive

cells around the scaffold wall struts embedded in fibrous soft tissues, indicating

interaction of activated macrophages with the scaffold material. These areas were

also found to highly express VEGF and stained positive for CD31 and vWF,

demonstrating presence of highly vascularised tissue (presence of large vessels also

confirmed in histological staining) with ongoing angiogenic signalling adjacent to

the scaffold wall and in close proximity to newly formed bone. This area also

showed ALP expression indicative of ongoing matrix mineralisation. Furthermore,

tissues in direct contact with scaffold surfaces (bone as well as soft tissue) showed

BMP 2&4-staining as marker for osteoinductive processes. We also found epitopes

for BMP2 to still be present in the remaining alginate islets (see Figure 29); however


no conclusion regarding its actual biological activity could be drawn from this.

Contrary to histological results from the rat femoral defect model [349], we found

small regions of enchondral ossification at the tip of callus advancing towards the

defect middle to be present in samples from group I and II (Safranin O / Fast Green

Stain). One possible explanation for this is the presence of largely hypoxic conditions

in such significant defect volumes, which are known to induce enchondral bone

formation rather than intramembranous ossification [160]. In samples from group III

however, only intramembranous ossification was found to be present (most likely

due to accelerated bone healing via BMP-growth factor-signalling and downstream

angiogenic signals). No alginate was found to be present outside the scaffold volume,

confirming the impression from the actual surgical implantation that porous scaffolds

spatially retained alginate hydrogels well.

In summary, we were able to confirm the high regenerative potential of the novel

spatiotemporal hybrid delivery system for rhBMP-2 in our preclinical ovine animal

tibial defect model. As hypothesized, the application led to significantly enhanced

bone formation and restoration of mechanical properties. Results obtained herein

directly parallel the results from the extensively investigated rat femoral defect

model. This is one of the rare cases were a tissue engineering application optimized

in a small animal model is apparently as efficient when directly applied in a

preclinical large animal model as well, bridging the scale-up gap between small

animal studies and potential clinical translation. We were able to further reduce the

relative dose of rhBMP-2 incorporated in the hydrogels by half (from 33.34µl/ml in

our previous rodent studies to 16.67µl/ml in this study), which is 90 times less the

current clinical dosage per ml. Normalizing the amount of rhBMP-2 by body weight

(assuming average body weight of 50kg per sheep), we delivered 0.02mg/kg

(rhBMP-2 weight/body weight) per sheep. This was equal to the dosage per body

weight used in the rat model and an approximate 7-fold reduction compared to the

current clinical dosage estimated as 0.136 mg/kg [362]. Since adverse side effects of

BMP-application have been potentially linked to supraphysiological quantities of

BMP delivered locally, this reduction may further improve safety in clinical

applications.

As outlined in the introduction, currently available evidence on the role of barrier

membranes for guided bone regeneration and restoration of large bone defects



derived from preclinical large animal studies is very rare [308]. Furthermore, only a

few clinical case series exist for the application in humans [363, 364]. The current

study addresses this shortcoming providing promising preclinical results

reproducible in two well-characterized small and large animal models including

long-term observations. Both the nano- or microfiber-meshes used as barrier

membranes and the functionalized alginate hydrogel matrix as well as the rhBMP-2

release kinetics have been extensively investigated and are subject of ongoing

research as well.

So far, ovine tibial defect studies by Gerber et al [312] and Gugula et al [313, 314]

using bioresorbable barrier membranes for guided bone regeneration showed that the

presence of autologous bone graft material (with or without addition of bone

substitute material) in the defect site was required to heal the segmental bone defects.

However, it has to be noted that the bone defects used by Gerber and Gugula of

greater length than in our ovine animal model (3cm vs. 7cm and 4cm, respectively)

and no external growth factors were added. Adipose-derived mesenchymal stem cells

(ADMSCs) transduced with the adenoviral vectors AdBMP2/AdBMP7 and

embedded in demineralized bone matrix (DBM) have been tested in a 10mm bone

distraction model for tibial fractures in sheep [365]. The study found complete bone

healing was achieved radiologically in less time (7–10 weeks) unlike other

experimental groups where consolidation was not achieved and bone deformation

was observed. However, the defect used in this study was rather small. As outlined

above, we have ourselves tested a PLGA-microparticle-based approach for sustained

delivery of rhBMP-2 (alone or combined with VEGF and PDGF) in a comparable

dosage of 1.12mg per sheep combined with cylindrical mPCL-scaffolds

manufactured by melt extrusion [359]. Although the 3cm tibial defects were bridged

in the BMP-2-group as well as the BMP-2/VEGF/PDGF-group in this study, total

bone volumes and biomechanical properties were significantly lower than in the

current study. This indicates that the efficacy of delivered rhBMP-2 is dependent on

the mode of delivery as well as on the scaffold type used.

RGD-functionalized alginate hydrogels provide not only a sustained release of

rhBMP-2 over time (in previous studies we conservatively estimated that 10% of the

encapsulated rhBMP-2 remained attached to alginate after 3 weeks [349]). They also

act as a matrix for ingrowing cells and are at the same time faster degrading than our

previously used cylindrical mPCL-scaffolds. The combination of such desirable


tissue engineering characteristics may be one of the key factors for the success of this

novel approach. Furthermore, the use of highly porous, functionalized microfiber

mesh-membranes as third generation tubular scaffolds to spatially retain hydrogels

but also applying guided bone regeneration-techniques is another important

contributor.

Our study is, to the best of our knowledge, the first ovine animal study to show that

the application of a spatiotemporal hybrid delivery system for rhBMP-2 utilizing

guided bone regeneration techniques and extended release of rhBMP-2 can

regenerate critical sized tibial segmental defects even without the addition of

autograft bone material. Our hybrid delivery system thereby offers a potential “off

the shelf”-solution in a single-staged procedure (and without a second surgical site

for bone graft harvesting) to overcome current limitations of clinical bone

regeneration strategies.

In conclusion, our results indicate that a spatially as well as temporally controlled

delivery strategy for osteogenic proteins such as rhBMP-2 (or other growth factors)

is a promising approach to enhance bone regeneration and overcome current

limitations in clinical applications. Using this large animal model we have herein

created preclinical evidence regarding feasibility and efficacy of this novel tissue

engineering strategy. Our large animal model results confirmed the results of

previous extensive investigations using a rodent animal model. Not only were we

able to upscale and transfer the methods to be applicable in an ovine large animal

model, we were also able to further reduce the dose of rhBMP-2/ml and apply a

consistently low 7-fold reduced rhBMP-2 dosage per kg body weight compared to

current clinical practice.

Following a linear scientific research model [67] (progression of research from initial

development of a material/application to in vitro testing, following by small animal

testing with consecutive large animal testing and then clinical testing) we have now

taken another significant step towards a potential clinical application. However,

further in studies (including but not limited to delivery of other growth factors via the

hybrid system, combination with other bone graft substitutes, further optimisation of

macro- and microstructure of tubular scaffold membranes and so forth) will be

necessary before clinical studies and ultimate translation from bench to bedside can

be expected.



3.5 ACKNOWLEDGEMENTS

The author would like to gratefully acknowledge all members of the QUT

histology laboratory and QUT BTM group (especially Flavia Medeiros Savi, Felicity

Lawrence and A/Prof Mia Woodruff) for their great work regarding the histological

and immunohistochemical analyses performed for this study.

Chapter 4: Establishment of a preclinical ovine animal model for the treatment of large volume 6cm-tibial

segmental defects 99

Chapter 4: Establishment of a preclinical

ovine animal model for the

treatment of large volume 6cm-

tibial segmental defects

4.1 INTRODUCTION

The treatment of large volume segmental bone defects (especially of load

bearing bones) and resulting non-unions is still considered a major challenge for

clinicians from various disciplines, including orthopaedic and plastic surgery [9, 23,

357]. This is especially true for the tibia (shinbone), which is the most commonly

fractured long bone in humans [5]. The average age of patients with tibial shaft

fractures is approx. 40 years, with teenage males being reported to have the highest

incidence [6, 7]. Treatment of (open) tibial fractures is often complex and poses a

significant risk of associated complications such as infection and non-union [8-10].

They are often associated with significant loss of bone substance and severe damage

to the surrounding soft tissue and carry a high risk for infection [12-14]. The average

time to union for uncomplicated tibial (shaft) fractures is approx. one year, but

complex cases can be much longer and require multiple surgical interventions [15].

Delayed union of bone or development of pseudarthrosis (non-union of bone) is

found in averagely 13% of all tibial fractures [16]. However, studies have reported

much higher non-union rates of up to 50-80% depending on the injury type, presence

of infection and surgical treatment [10, 16]. The consequences of suffering a severe

(open) tibial fracture with threatening limb loss, potential consecutive delayed bone

healing or development of pseudarthrosis can be devastating for patients, their

families/social environment and the entire society (loss of productivity, health care

cost etc.). This is drastically illustrated by the fact that only 28% of patients suffering

severe open tibial fractures resume full function and are able return to their previous

employment [1]. Non-unions of tibial shaft fractures are associated with substantial

healthcare resource use, common and prolonged use of strong opioids, multiple

revision surgeries and high per-patient costs [18, 19] and represent a significant

100Chapter 4: Establishment of a preclinical ovine animal model for the treatment of large volume 6cm-tibial segmental defects

burden for the individual patient as well as for the healthcare system and society in

general.

Given the limitations of current clinical treatment options, there is a clear

demand for novel treatment alternatives offering “off-the-shelf”-solutions (ideally in

a single-stage procedure) for these challenging bone defects. Consequently, the field

of bone tissue engineering has received considerable interest from the surgical

community over the last years and a number of promising approaches are being

investigated [45, 366]. However, the number of tissue engineering approaches that

have actually reached clinical application yet is very limited compared to the

plethora of techniques investigated in the laboratories [65]. While segmental long

bone defects are often studied in small animal models (especially rodents) first [367,

368], scaling-up to large animal models for clinical modelling and efficacy

prediction is necessary for translation of the tissue engineering product from bench to

bedside. Mass and volume challenges for scaffold-based tissue engineering

encountered in large animals as well surgical fixation techniques closely resembling

the clinical situation cannot be tested either in vitro or in small animal models [65].

Despite several disadvantages (increased costs, long life spans, low bone turnover

rates, difficulties in standardization and so on) pre-clinical trials in large animals,

with orthotopic implantation sites, with relevant loading conditions and with similar

surgical techniques as used in the final procedure in humans are necessary [260].

Sheep are frequently used in orthopaedic large animal studies because they limb

loading conditions similar to humans, are of a similar body weight as humans and

feature long bone dimension suitable for the application of human implants [369].

Furthermore, aged sheep (>6-7 years) display secondary, Haversian (osteonal)

remodelling which is the predominant mode of bone remodelling in humans.

Sheep animal models are being used to investigate bone healing after tibial

osteotomies with different fixation techniques, varying mechanical conditions,

combined with growth factors or bone substitutes/bone grafts as well as with or

without present surgical site infection [see for example [370-376]]. Cortical or

cancellous (most often circular) tibial defects filled with various tissue engineering

constructs are also commonly used in sheep animal models [377, 378]. Reviewing

the literature, several sheep models utilizing segmental tibial defects of sizes ranging



from 1.8-3.5cm stabilized with intramedullary nailing [379-382], external fixation

[383-386] or (single or double) plating techniques [387-391] can be found.

However, there are only a handful of published studies on ovine animal models

with defects greater than 4cm in the tibial diaphysis: Christou et al. have established

a 5-cm mid-diaphyseal osteoperiosteal tibial defect in aged sheep (> 5 years) that was

stabilized by using an 8-mm stainless-steel cross-locked reamed intramedullary nail

[392, 393]. Another study by Pluhar and colleagues used 5cm diaphysial defects in

tibiae of ‘skeletally mature’ sheep stabilised with reamed intramedullary interlocking

nails [394]. Mastrogiacomo et al reported an ovine animal model creating a 4.8cm

segmental defect in the mid-third diaphysis of 2-year old ewes stabilized with a

single 4mm neutralizing plate) [395]. Gogolewski and colleagues published data on

guided bone regeneration techniques investigated in a 4cm long osteoperiosteal tibial

defect in swiss mountain sheep (age not reported) stabilized with a bilateral AO

external fixator [313, 314]. In two other studies a 7cm diaphysial tibial defect

stabilised with an unreamed locked intramedullary 8mm-nail in ‘adult’ sheep was

investigated [312, 396].

Most of the above mentioned large segmental tibial defect studies used reamed

intramedullary locking nails as fixation devices. While an intramedullary force-

conduction provides increased stability at the defect site in long bones,

intramedullary reaming is known to alter cortical blood supply and furthermore the

implant is a foreign body inside the defect area that could potentially influence or

impede (endosteal) healing patterns. Additionally, only one [384] of the above

mentioned studies did include long term results up to 24 months after implantation.

Furthermore, except for the work by Christou et al. [392, 393], studies used relatively

young sheep, in which bone microstructure is known to be significantly different

from humans (predominantly primary bone structure in comparison with largely

secondary bone of humans) [369].

Over the last decade, out research group has established and extensively

characterised a preclinical 3cm tibial defect ovine animal model at Queensland

University of Technology (QUT) [71-78, 361, 397, 398]. We have trialled a number

of different tissue engineering applications using this model as a testbed providing

highly standardized experimental protocols and well established control groups [72-

76, 78, 359, 361]. We were thereby able to create statistically significant preclinical


evidence on the regenerative potential of different treatment strategies enabling

amongst others direct comparisons between different treatment groups and tissue

engineering applications. However, bone defects encountered in clinical practice are

often of larger volume and even more challenging than the conditions we created in

our 3cm tibial defect model. Next to trauma and fracture non-unions as a cause of

bone substance loss, this is especially true for bone defects due to revision surgery

after failed arthroplasties as well as in orthopaedic oncology [128]; and numbers are

predicted to grow in our aging population. We therefore aimed to further optimize

the experimental setting used to investigate tissue engineering applications in order

to reflect the current clinical situation.

In this study we report the establishment of a large volume 6cm tibial segmental

defect model based on our experiences and expertise gained from the well-

established and characterized 3cm tibial defect ovine anima model at QUT. We have

successfully adapted study protocol, surgical technique as well as our analyses

standards to be applicable in this larger sized ovine tibial defect model as well. In

order to minimize confounding variables in the pilot study, we used a porous medical

grade Polycaprolactone-Tricalciumphosphate (mPCL-TCP)-scaffold combined with

Bone Morphogenetic Protein-7 (rhBMP-7) and Platelet-Rich- Plasma (PRP) as

previously used to successfully regenerate 3cm tibial defects in our ovine model (of

which results have been extensively analysed) [75, 78, 361, 397, 398].



4.2 MATERIAL AND METHODS

4.2.1 Animal Ethics Approval and Code of Practice

Prior to surgeries, ethical approval was gained from the University Animal Ethics

Committee (UAEC) of the Queensland University of Technology, Brisbane,

Australia (Ethics Approval Number 1000001139). All animal surgeries were

performed at the QUT Medical Engineering Research Facility (MERF), The Prince

Charles Hospital, Chermside, Brisbane, QLD, Australia. The study was conducted in

accordance with all requirements of the Australian Code of Practice for the Care and

Use of Animals for Scientific Purposes. Eight male Merino sheep (50-60 kilogram

bodyweight at day of surgery, age ≥6 years) were included in the study with two

different end points (3 months and 12 months post-surgery, n=4 respectively).

4.2.2 Scaffold design and fabrication

Three-dimensional porous biodegradable scaffolds were fabricated by fused

deposition modelling (FDM) consisting of medical grade polycaprolactone (mPCL,

80 wt.%) and β-tricalcium phosphate (TCP, 20 wt.%) (outer diameter 20 mm, inner

diameter 9 mm, height 60 mm) (Osteopore International, Singapore). Scaffold

dimension were derived from radiographic analysis of anatomical dimensions of 10

sheep tibiae. Structural parameters were set by Computer Aided Design (CAD)

resulting in 70% porosity with 100% pore interconnectivity and a pore size of 350-

500 μm. Filaments (300 μm diameter) were deposited in a 0/90° pattern with a

separation of 1200 μm. Nine linear holes (diameter 2mm) were punched into the

scaffold (using a biopsy punch) after fabrication to promote ingrowth of larger sized

blood vessels after planned implantation in close proximity to the neurovascular

bundle. Prior to implantation, scaffolds were surface treated with 1 M NaOH for 6

hours and washed five times with phosphate-buffered saline (PBS) to render the

scaffold surface more hydrophilic. Scaffolds were then sterilised by incubation in

70% ethanol for 5 min followed by complete evaporation and subsequent UV

irradiation for 60 min.


4.2.3 Characterisation of Mechanical properties of 6cm mPCL_TCP Scaffolds

and Scaffold construct biomechanical testing

This work was conducted by Dr. Stephanie Fountain during her PhD research.

Experimental protocols as well as results have been published elsewhere [399]. The

author of this thesis assisted with the mechanical testing of the scaffolds.

4.2.4 Preparation of Platelet Rich Plasma (PRP) and loading of Scaffolds with

rhBMP-7

Autologous platelet-rich plasma (PRP) was obtained from each sheep by collecting

80 ml of peripheral venous blood from the external jugular vein in 3.5 ml monovettes

containing sodium citrate (3.8%) at a ratio of 9 volumes of blood to 1 volume of

sodium citrate [400]. The citrated blood was then transferred to eight 15ml Falcon

tubes and centrifuged in a standard laboratory centrifuge for 20 min at 2400 r.p.m.

(580g). Next the yellow plasma layer was transferred to a fresh 50ml Falcon tube and

the platelets pelleted in a second centrifugation step at 3600 r.p.m. for 10 min

(1300g) [401]. The pellet was then resuspended in 2 ml of plasma to form PRP. After

sterilisation, scaffolds were placed in large petri dishes and loaded with 2mg rhBMP-

7 (Olympus Biotech Corporation) suspended in 2ml autologous PRP. Afterwards

PRP was clotted with thrombin (5 U ml) to contain PRP and rhBMP-7 in the

scaffold.

4.2.5 Surgical procedure

Based on the expertise from our existing 3cm ovine tibial defect model [72-75, 78,

361], the surgical technique was modified in order to account for the larger 6cm

segmental tibial defect (Figure 31): Eight male Merino sheep (50-60 kilogram

bodyweight at day of surgery, age ≥6 years) were included in the study. All animals

were placed in right lateral position with the right hind leg exposed. Antimicrobial

washing was performed using a Chlorhexidine Solution followed by sterile surgical

draping of the hind leg. Opsite Incise Drape (Smith&Nephew, USA) was applied t

cover the surgical site. The tibial bone was approach by a longitudinal approx. 15cm

long skin incision on the medial aspect of the tibial diaphysis (Figure 31, A).



Underlying soft tissue was split and the tibia exposed. A 12 hole broad (5.6mm

thickness) Dynamic Compression Plate (DCP, Synthes®) was inserted on top of the

medial tibial surface, after being bent to fit the anatomical shape of the tibia if

necessary. The distal end of the plate was placed 1cm proximal (above) the medial

malleolus (ankle) to ensure exact and standardised plate placement and defect

position. Four pilot holes for later screw placement were drilled proximally and

distally, respectively, using the four outmost holes on each side of the plate as

template (Figure 31, B). The DCP was temporarily fixed with two screws (Figure 31,

C) and the middle of the plate was determined and marked with an incision into the

periosteum. This incision equalled the middle of the bone defect to be created in the

following. Screws and DCP were removed, a distance of 3cm was measured towards

each side proximally and distally from the defined defect centre and these osteotomy

sites marked down (Figure 31, D). Soft tissue inserting in the designated defect area

was carefully detached from the bone to avoid muscle, nerve and blood vessel

damage. The periosteum was then opened and carefully detached from the bone in

the defect area. Parallel osteotomies at the preassigned defect margins were

performed using an oscillating saw perpendicular to the tibial longitudinal axis under

constant irrigation with saline solution to prevent heat-induced osteonecrosis (Figure

31, E). The cut out 6cm bone segment was then removed and the periosteum

completely removed in the defect site (Figure 31, F-G). Special care was taken to

remove the thick periosteal strand adjacent to the lateral neurovascular bundle

(Figure 31, H-I, Asterisk marks neurovascular bundle). Former studies have shown

that the defect heals with treatment (= is not a critical sized defect), if this part of the

periosteum remains in the defect site. To prevent endogenous regeneration from

adjacent periosteal tissue, the remaining distal and proximal bone segments were

additionally circularly denuded of periosteum over a length of 1cm (Figure 31, J).

The DCP was then firmly fixed to the proximal tibial bone segment with 4 screws

(Figure 31, K). Afterwards, the bone segments were realigned and the DCP loosely

fixed to the distal segment with 4 screws (Figure 31, L-M). The 6cm mPCL-TCP

scaffold loaded with PRP and 2mg rhBMP-7 (as described above) was then press-

fitted into the defect site and screws in the distal tibial segment tightened to

definitively stabilise the defect site (Figure 31, N-O). The wound was closed in

layers, using 2-0 Monocryl (Ethicon©) for soft tissue/subcutaneous tissue and 3-0

Novafil (Syneture©) sutures for skin closure. Antibiotic spray and sterile dressing


was applied to the surgical site before X-ray analysis to confirm correct implant

placement. Afterwards, a circular cast (Vet-Lite©, D.L.C Australia) was applied to

the right hind leg before the animal was recovered from anaesthesia. Sheep were kept

indoors for three weeks while in cast to limit physical activity. Three weeks after

surgery, all cast material was removed and the sheep kept indoors for another week

for close observation. The sheep were then released into the paddocks at MERF and

transported back to the animal agistment facility afterwards.

Figure 31: Surgical Technique for 6cm tibial defect animal model with mPCL-TCP scaffold +

PRP + 2mg rhBMP-7. For details please refer to description in text. Asterisk in I = Neurovascular

bundle.

Since our study group has already proven the critical-sized nature of a 3cm tibial

defect created with similar surgical techniques, we did not include an empty defect-

control group in this study. Increasing the defect size by 100% will not lead to an

enhanced healing potential compared to the smaller 3cm defect. Therefore, a 6cm

tibial defect created in the same way as the 3m defect previously can safely be

considered a critical-sized defect. Consequently, the number of animals needed for

the study could be reduced in order to account for best possible practice and animal

welfare.

4.2.6 Conventional X-ray analysis

Conventional X-rays in two standard planes (anterior–posterior and medial–lateral)

were taken after surgical procedure to confirm correct implant (plate, screws)



placement and scaffold positioning. Serial conventional X-ray analyses (3.2 mAs,

65kV; Philips, Australia) in two standard planes (anterior–posterior and medial–

lateral) were performed at 3, 6, 9 and 12 months post-surgery to evaluate formation

of newly formed mineralised bone tissue and time point of defect bridging.

4.2.7 Euthanasia of sheep and harvesting of specimens

Humane euthanasia was performed at 3months post-surgery (n=4, group I) and 12

months post-surgery (n=4, group II), respectively, by intravenous injection of 60

mg/kg pentobarbital sodium [Lethabarb; Virbac Animal Health, Milperra, New

South Wales, Australia, http://www.virbac.com)]. Both hind legs (operated

experimental right tibia as well as non-operated control left tibia) were retrieved

from each sheep for further analysis. Surrounding musculature and soft tissue were

carefully removed without damaging the defect area. Specimens were then frozen

down at -20°C prior to analysis.

4.2.8 Biomechanical testing

All surgical implants (DCP and screws) were completely removed from experimental

tibiae prior to biomechanical testing taking care not to damage the defect area. Bone

ends were then fixed in custom-made jigs using Paladur (Heraeus Kulzer) dental

acrylic. After hardening of the Paladur, samples were mounted in a biaxial testing

machine (Instron 8874, Instron, Norwood, USA). Torsion testing was conducted

under angular displacement control at an angular velocity of 0.5°/s and a constant

compressive preload of 0.05 kN until first signs of failure occurred. Total Maximum

torsional moment (TM) and torsional stiffness (TS) values were calculated and

normalized against the measured values of the contralateral, non-operated tibia of the

same animal. Detailed protocols for biomechanical testing can be found in [75].

Following biomechanical testing all experimental right tibial specimens were cut to a

total length of 8cm (complete 6cm defect site length plus 1cm adjacent host bone on

each end).


4.2.9 Micro computed tomography (µCT)

After mechanical testing microCT scans of the defect site and adjacent host bone

were performed using standardized protocols as published by our group before [75].

All samples were placed in custom-made tubes and scanned in a Viva40© μCT

(Scanco Medical AG, Brüttisellen, Switzerland) with a voxel size of 36 μm. The X-

ray tube was operated at 55 kV and 145 μA. A threshold of 210 HU (519.2 mg of

hydroxyapatite per cubic centimetre), a Gaussian filter width of 0.8 and filter support

of 1.0 were chosen to best analyse the morphology of mineralised tissue and to

exclude scaffold and soft tissue [402]. Bone volume (BV) [mm3] within the defect

was calculated using the supplied manufacturer software. The analysed volume of

interest (VOI) included the defect region and adjacent host bone only. Total bone

volume (TBV) was measured for the complete defect volume. In analogy to mCT

analyses performed on our 3cm tibial defect samples in previous studies [72-75, 78,

361], we furthermore analyses axial and radial bone volume (BV) distribution. For

axial BV distribution the total length of the defect was divided into three parts of

equal length (proximal, middle, distal; 2cm length each). Radial bone distribution

was described by defining three volumes of interest (VOI): Scaffold inner duct

(VOIinner_duct), scaffold wall (VOIscaffold – VOIinner_duct), and scaffold periphery

(VOItotal – VOIscaffold).

4.2.10 Histology and Immunohistochemistry

Sample processing as well as histological and immunohistochemical staining was

performed by Flavia Medeiros Savi assisted by the author and under supervision of

A/Prof. Mia Woodruff. Analysis of the results was performed by the author in

collaboration with the above mentioned persons. After biomechanical testing and

microCT analyses, samples were cut to a length of 6 cm (defect size) plus an

additional 3mm of proximal and distal host bone (longer portions of host bone would

have increased total sample length above maximum available length of histology

slides, making an analyses of a sample in toto impossible). For histological analysis,

the samples were then sectioned in transverse and sagittal planes (please refer to

cutting schematics shown in figures below for further details). Samples were then

fixed in 10% neutral buffered formalin for 1 week. For paraffin format the transverse



planes were cut into three sections (P1, P2 and P3). The bone samples were

decalcified in 10% EDTA for 6-8 weeks at 37°C using a rapid decalcifier at an input

voltage 230V-59Hz, 8A and 450rpm (Kos Milestone microwave model 67051,

ABACUS, Brisbane, Australia). The samples were then serially dehydrated in

ethanol in a tissue processor (Excelsior ES, Thermo Scientific, Franklin, MA, USA),

and embedded in in molten paraffin wax at 60C (Thermo Shandon Histocentre 3

Embedding Station, Thermo Scientific, Brisbane, Australia). 10 sections were cut at

5 µm with a Leica RM2235 rotary microtome (Leica Biosystems, Nussloch

Germany). Paraffin ribbons were flattened on a water bath (Labec, Marrickville,

Australia) at 40C and collected onto polysine microscope slides (Thermo Scientific,

Brisbane, Australia) prior to drying at 60C for 16 h. Two slides were then stained

with Hematoxylin and Eosin staining (HD scientific, Wetherill Park, Australia) &

Eosin (HD scientific, Wetherill Park, Australia) using a Leica Autostainer XL (Leica

Biosystems, Nussloch, Germany). The slides were scanned using a Leica SCN 400

slide scanner (Leica Microsystems, Wetzlar Germany) with a 20x objective. For

resin format samples were sectioned in two sagittal planes (R1 and R2) at 2mm thick

slices using an EXAKT 310 Diamond Band Saw (EXAKT Apparatebau GmbH &

Co.KG, Norderstedt, Germany). Following degreasing with xylene, the allocated

samples for Technovit 9100 New® were processed and embedded in the low-

temperature embedding system Technovit 9100 New® (Heraeus Kulzer GmbH,

Germany). For ground section format the mounted resin blocks were sectioned

longitudinally at 200µm using a EXAKT 310 Diamond Band Saw and subsequently

ground at 50µm using a EXAKT 400CS micro grinder (EXAKT Apparatebau GmbH

& Co.KG, Norderstedt, Germany) according to the technique described in Donath

1995. Histological assessment was performed using Goldner’s trichrome staining.

For immunohistochemistry, paraffin sections were deparaffinised with xylene and

rehydrated with serial concentrations of ethanol. Subsequently, sections were rinsed

in distilled water and placed in 0.2 M Tris-HCl buffer (pH 7.4). Endogenous

peroxidase activity was blocked by incubating the sections in 3% H2O2 in Tris-HCl

for 20 min. This was followed by three washes with Tris buffer (pH 7.4) for 2 min

each. Sections were incubated with Proteinase K (DAKO, Botany, Australia) for 20

min and subsequently incubated with 2% bovine serum albumin (BSA) (Sigma,

Sydney, Australia) in DAKO antibody diluent (DAKO) in a humidified chamber at

room temperature for 60 min to block non-specific binding sites. Afterwards,


immunohistochemical staining was performed using primary antibodies specific to

the osteogenic markers:

1. Type I collagen Ab 34710 dilution 1:100 Dab: 2 min (rabbit polyclonal

Abcam, Cambridge, UK).

2. Bone Morphogenic factor 2&4 SC 137087 dilution 1:50 Dab: 1:30 min

(Santa Cruz, Biotechnology, CA, USA ).

3. Von Willebrand Factor A0082 dilution 1:700 Dab: 1:30min (Ready to use,

rabbit polyclonal, IR527, Dako, Glostrup, Denmark)

4. Cluster of differentiation 31(CD31) SC 1506R dilution 1:1000 Dab: 1:30min

(Santa Cruz, Biotechnology, CA, USA).

5. Cluster of differentiation 68 (CD68) Ab 125212 dilution 1:300 Dab: 2min

(Abcam, Cambridge, UK).

6. Alkaline phosphatase (ALP) Ab 108337 dilution 1:500 Dab: 5min (Abcam,

Cambridge, UK).

7. Vascular Endothelial growth factor (VEGF) SC 152 dilution 1:500 Dab:

1:30min (Santa Cruz, Biotechnology, CA, USA).

The sections were incubated with the specific antibody in humidified chambers at

4°C overnight. Sections were then washed three times for 2 min with Tris buffer (pH

7.4) and incubated with peroxidase labelled dextran polymer conjugated to goat anti-

mouse and anti-rabbit immunoglobulins (DAKO EnVision+ Dual Link System

Peroxidase, DAKO) at room temperature in humidified chambers for 60 min. Colour

was developed using a liquid 3,3-diaminobenzidine (DAB) based system (DAKO).

Kaiser’s glycerol gelatin (DAKO) was used for coverslip mounting.

4.2.11 Statistical analysis

Statistical analysis was performed using SigmaPlot statistical software (Systat-

Software Inc.) using Normality Test (Shapiro-Wilk) and, if passed, Equal Variance

Test (Brown-Forsythe). If Normality Test failed, Kruskal-Wallis One Way Analysis

of Variance on Ranks was performed. If Equal Variance Test failed, Mann-Whitney

Rank Sum Test was performed instead. Differences between groups were considered

to be statistically significant at p values <0.05.



4.3 RESULTS

4.3.1 Surgical procedure and postoperative follow-up period

All 8 sheep (n=4 animals per group) tolerated the surgical procedure well.

Postoperative X-rays (anteroposterior and mediolateral standard planes) confirmed

the correct position of the diaphyseal defect site as well regular positioning of the

DCP and screws. Postoperative recovery was achieved without any complications in

all sheep. All animals included in this study were in good health and survived the

experimental period, gaining weight in the months following surgery. Postoperative

follow-up was without adverse events except for an implant failure (breakage of

DCP in the middle of the plate located next to the defect middle) in one sheep on the

day of planned endpoint euthanasia 12 months after surgery. On necropsy and

radiographic examination, the defect site showed substantial new bone formation

extending from both ends inwards, but with a non-union site in the middle of the

defect. It can be speculated that the DC plate broke as a result of chronic fatigue due

to defect non-union combined with supra-physiologically high peak loads

experienced during the transport from the animal agistment facility back to MERF.

Slight bending of the DCP at the distal end of the defect site was noticed in one

sheep. However, the defect site showed bony bridging with hypertrophic callus

formation at 6 months post-surgery and no progressive DCP bending or implant

failure occurred. Another sheep had a breakage of a single screw (most cranial screw

in tibial bone) without further consequences.

4.3.2 Conventional X-ray analysis

Serial X-ray analyses in two standard planes (anteroposterior and mediolateral) were

performed at 3, 6, 9 and 12 months post-surgery in group II (n=4).

All animals in group I (n=4) were humanely euthanised after conventional X-ray

analysis at 3 months post-surgery. X-ray-images at 3 months after surgery (Figure 32

A and Figure 33 A, left row) revealed substantial new bone formation to be present

in the defect site in all animals. New bone was found to be extending from both bone

ends growing into the scaffold in the defect area, with the proximal defect site

showing increased bone formation (a healing pattern observed in previous studies as


well, most likely due to soft tissue coverage). The scaffold structure was visualized

(as void) in between the regenerating bone confirming structural guidance of the

porous mPCL-TCP scaffolds. At three months after surgery, the defect site was

found to be bridged radiologically in 62.5% (5/8) of all animals included in the study

(3 months-group and 12 months-group taken together), while the rest displayed new

bone formation in the defect site close to bridging (25%, 2 of 8 sheep) or with at least

50% of the defect length bridged (12,5%, 1 out of eight sheep). At 3 month after

surgery, radiopaque mineralized bone was mainly seen along/inside the exterior parts

of the scaffold (periosteal bone formation) or inside the scaffolds’ internal lumen,

while the porous scaffold wall itself was only partially invaded by mineralized bone.

Serial x-rays of animals in group II (Figure 33, A) showed progressive bone

formation in the defect site over the following 9 months with bony defect bridging

present at 12 months in 3 out of 4 animals (75%). Bone formation along the scaffold

structure and (compared to the three month group) now also growing into the

scaffold wall was evident, with the scaffolds appearing as non-mineralized void in

the defect volume. In one sheep, non-union of the bone defect persisted in the middle

part of the defect site. As outlined above, this sheep experienced implant failure with

breakage of the DCP at the day of euthanasia. Furthermore, breakage of a single

screw (most cranial screw) was observed at 3 months in one sheep of group II

without dislocation of the entire construct or progressive screw dislocation in the

following months. No hypertrophic callus formation was seen in this sheep.



Figure 32: Conventional X-ray images and 3D-reconstructions of mineralized tissue from µCT

in 3-months-group (group I). (A) X-ray images in two standard planes of specimen at 3-months after

surgery. Substantial new bone formation was seen in the defect site, with 75% (3 out of 4 sheep)

showing bone defect bridging at this early time point. Red arrows indicated samples of which

microcomputed tomography scans are shown below. (B) Representative images of 3d-reconstructions

of mineralized tissue from microcomputed tomography (µCT) analyses. Left image shows 3 Phase-

Segmentation of old (white) and new (grey) bone, Scaffold and void are black . Right image shows

Microporosity-Segmentation of both old and new bone. Scaffold and void are black. Solid bone is

white and the darker the grey level the more porous the bone (analysis done at ANU Canberra) (C)

Representative images of 3d-reconstructions of mineralized tissue from µCT analysis in two planes.

Both B and C confirm mainly periosteal new bone formation along the scaffolds outer surface and

inside outermost parts of the scaffold wall to be present.


Figure 33: Conventional X-ray images and 3D-reconstructions of mineralized tissue from mCT

in 12 month-group (group II). (A) Serial X-ray images (a.p.-plane) at 3, 6, 9 and 12 months for all

animals in group II. Progressive new bone formation in the defect volume over the course of the

study. At 12 months, 3 of 4 animals showed bon bridging of the defect site (for sheep II, bridging was

confirmed in X-rays of lateral plane as well as µCT). However, in one animal (sheep IV, bottom row)

non-union persisted and plate breakage was experienced at 12 months after surgery (Asterisk indicates

plate breakage site). Failure of a single screw (red arrow head top row) was seen at 3 months in one

animal without further dislocation over time. Red arrows indicate specimens of which µCT-3d

reconstructions are shown on b and C. (B) Representative images of 3d-reconstructions of mineralized

tissue from µCT analysis in two planes. Compared to the 3 months-time point, substantial bone

formation inside the scaffold wall as well as the endosteal scaffold lumen was now present. The

scaffold was fully integrated into the newly formed bone and differentiation into cortex (including

scaffold wall) and a medullary canal (inside the internal lumen of the scaffold) was visible. (C)

Representative images of 3d-reconstructions of mineralized tissue from µCT analysis in two planes of

non-union specimen. In contrast to the other samples of group II, new bone is mainly formed inside

and along the internal scaffold lumen. Little bone ingrowth into the porous scaffold wall or around the

outer scaffold surface was visible.



4.3.3 Micro computed tomography (µCT)

3D-reconstruction of mineralized tissue for data obtained from MicroCT scans at

three month (Figure 32, B and C) confirmed bony defect bridging to be present in 3

out of 4 samples (75%) as already assessed on conventional X-ray analysis. New

formation of mineralized tissue was found to have taken place mainly at the interface

scaffold/host bone as well as along the outside surface of the scaffold and in the outer

parts of the porous scaffold wall. The internal lumen of the scaffold as well as

internal parts of the scaffold wall were not filled with mineralized tissue yet, except

for the interface areas directly adjacent to the host bone proximally and distally.

Radial bone volume distribution confirmed newly formed bone to be mainly present

around the scaffold externally as well as in the scaffold wall.

In accordance with visual assessment of 3D-reconstructions from mCT-scans,

statistical analysis (Figure 34) of radial bone volume (BV) distribution at three

months confirmed significantly higher (p<0.05) bone volumes to be present around

the scaffold externally compared to the inside lumen of the scaffold. A trend towards

higher BV inside the scaffold wall compared to the internal scaffold lumen was

observed, but failed to reach statistical significance (p=0.079). No significant

differences for axial bone volume distribution were found, though a trend towards

increased BV in the proximal as well as the distal third of the defect was found (bone

extending from both defect ends inwards).

At 12 months after surgery, ex vivo µCT analyses (Figure 33, B and C) showed bony

defect bridging in 3 of 4 sheep (75%) and one persisting bone non-union (25%).

Compared to the 3 months-time point, new bone formation inside the scaffold wall as

well as the endosteal scaffold lumen was now also found to be present for the

majority of samples. The scaffold was fully integrated into the newly formed bone

and differentiation into cortex (including scaffold wall) and a medullary canal (inside

the internal lumen of the scaffold) was visible. In contrast to other samples of group

II, the non-union specimen (Figure 33, C) exhibited new bone formation mainly

inside the internal scaffold lumen extending towards the defect middle. Little bone

ingrowth into the porous scaffold wall or around the outer scaffold surface was

visible. No statistically significant differences regarding axial or radial bone volume

distribution were found in group II at 12 months after surgery (p=0,560 and p= 0,066

for axial and radial bone volume distribution, respectively) (Figure 35, A).


Comparing results of group I and group II (Figure 35, B), a trend for increased Total

Bone Volumes (TBV) after 12 months was observed, but no statistically significant

differences were found (p= 0,139). Axial bone volume distribution was not

significantly different between the two groups, either (p= 0,0891 for proximal third,

and p= 0,421 for middle third, and p= 0,0730 for distal third, respectively). In

accordance with the observations made on plain radiographs and µCT-images, we

found significantly increased bone volumes in the scaffolds inner duct at 12 months

compared to three months (p= 0,00391). A trend towards increased bone volumes in

the scaffold wall was also observed, but without statistical significance (p=0.1). No

significant differences between bone volumes in scaffold periphery were found

between group I and II (p= 0,761).

Figure 34: Microcomputed tomography-results at three months post-surgery. (A) Schematic

showing analysis of axial and radial bone volume distribution. (B) Results at three month after surgery

for radial and axial bone volume distribution. Significantly higher bone volumes in the scaffold

periphery compared to the scaffold inner duct were found (p<0.05, Asterisk). Although a trend

towards higher bone volumes in the scaffold wall was found, results were not significant (p=0.079).

Axial bone volume distribution showed trend towards higher bone volumes in proximal and distal

defect sites, but without statistical significance.



Figure 35: Statistical analysis of microcomputed tomography at 12 months post-surgery (A) and

comparison between group I and II (B). For details please refer to text. Asterisk indicates

statistically significant differences (p<0.05).


4.3.4 Biomechanical analysis

Biomechanical testing (Figure 36) was performed using the same protocols as

applied in our 3cm defect model [75]. Results at three months showed significantly

lower (both p ≤ 0.001) values for torsional moment (TM) and torsional stiffness (TS)

of the operated experimental tibiae compared to the contralateral, non-operated tibiae

of the same animals. At twelve months after surgery, overall results were still

significantly lower for TS (p= 0,039) and TM (p=0.0121) for experimental tibiae

compared to non-operated contralateral tibiae. An increase in total results for TS and

TM at twelve month compared to the 3 months-time point was observed, but failed to

reach statistical significance (p= 0,113 and p= 0,114 for TM and TS, respectively). A

great variability was observed in group II with one animal showing non-union, one

animal showing results in the range of 5% of contralateral side for TS and TM and

two animal reaching TS and TS-values of around 50% of the contralateral non-

operated tibia.

Figure 36: Statistical analysis of biomechanical testing. (A) Comparison of total values for Torsional

Stiffness (TS) and maximum Torque (TM) between operated right tibia and non-operated left control tibia

showed significantly lower values at three months (group I, n=4). At twelve months (n=3), TS and TM was

still significantly lower for operated legs compared to non-operated control tibiae of the same animal.

Asterisks indicate significant differences between groups (p<0.05). (B) Comparison between the two study

groups revealed no statistically significant differences between 3- and 12-months time-points when

normalized against results of contralateral non-operated tibiae.



4.3.5 Histology and Immunohistochemistry

Histological sections stained with Hematoxylin & Eosin (H&E, Paraffin-embedded

sections) as well as Goldner’s trichrome Stain (Resin-embedded sections) confirmed

mineralized bone tissue to be present mainly in the scaffold periphery and outer

scaffold wall at three months (Figure 33), while at 12 months (Figure 34) bone had

also grown into the entire scaffold wall and internal scaffold duct. The scaffold was

found to be well integrated into the newly formed bone with full osteointegration

present at the host bone/scaffold-interface. Bone structure was plexiform with woven

bone present and lamellar bone on its’ surface in the inner duct of the scaffold as

well as the scaffold periphery. Inside the scaffold wall more organized lamellar bone

was present. Mature osteocytes embedded in lacunae, osteoblasts depositing osteoid

and bone-resorbing osteoclasts were present. The newly formed bone was well

vascularized. Primary osteon formation with surrounded by interstitial matrix was

visible on the newly formed cortex area. At 12 months, Harversian remodelling was

found mainly in the cortex-area. Plexiform woven bone in the defect had been widely

replaced by higher organized lamellar bone structures around the scaffold struts.

New bone as well as adjacent soft tissues in the defect area stained

immunohistochemically positive for Collagen I as (early) osteogenic marker. At

three month, the interface of the advancing callus with the soft tissue was positive for

ALP. Soft tissue between the scaffold struts and inside the scaffolds internal duct

showed widespread staining for VEGF, indicating strong angiogenic signalling.

CD31 and vWF was also found in these areas with mature vessel formation. At

twelve month, ALP was less pronounced and could be found mainly at the interface

of newly formed bone with scaffold struts as well as in the peripheral cortex region

in close proximity to the periosteum. VEGF and CD 31 were located mainly in the

newly formed periosteal and endosteal regions. Inside the scaffold wall, positive

staining was less frequent and mainly visualized in close proximity to scaffold struts

or inside newly formed bone. Blood vessels of varying diameter staining positive for

vWF were visualized located in Harversian Canals as well as in soft tissue inside and

outside the scaffold wall. At three month, CD68-positive staining was found to be

widespread inside the scaffold wall and in adjacent soft tissues. CD68-staining was

less frequent at 12 months and mainly in close proximity to scaffold struts. While

BMP was present in soft tissue as well as the interface to newly formed bone inside

the defect at three months, BMP-positive staining was located mainly in close


proximity to scaffold struts as well as sub-periosteal at twelve months. The non-

union specimen (at 12 months after surgery) displayed features as described above

for three months samples with widespread ALP-, Collagen I- and VEGF-positive

staining indicative of ongoing remodelling processes.

Figure 37: Overview of results from histological stains and immunohistochemical analysis of

group I (3 months time-point, representative sample specimen). Top left row shows X-ray images

at time of sample harvesting with schematic of sample explantation and processing. Sagittal plane 3D-

reconstruction of microcomputed tomography of corresponding sample shows amount of mineralized

tissue with defect margins. Images from undecalcified resin-embedded sagittal sections (as indicated

in schematic) stained with Goldner’s trichrome are shown on top right Bottom row left shows

schematic of horizontal sample cuts for decalcification and Paraffin embedding. Representative

images for all three defect regions stained with Haematoxylin Eosin (H&E) as well as

immunohistochemical analyses (antibody against epitopes listed on top of each column) are shown in

the bottom row. Black bar indicates 100µm for H&E as well as IHC images. Figure designed by

Flavia Medeiros Savi and the author.



Figure 38: Overview of results from histological stains and immunohistochemical analysis of

group II (12 months time-point, representative sample specimen). Top left row shows X-ray

images at time of sample harvesting with schematic of sample explantation and processing. Sagittal

plane 3D-reconstruction of microcomputed tomography of corresponding sample shows amount of

mineralized tissue with defect margins. Images from undecalcified resin-embedded sagittal sections

(as indicated in schematic) stained with Goldner’s trichrome are shown on top right Bottom row left

shows schematic of horizontal sample cuts for decalcification and Paraffin embedding. Representative

images for all three defect regions stained with Haematoxylin Eosin (H&E) as well as

immunohistochemical analyses (antibody against epitopes listed on top of each column) are shown in

the bottom row. Black bar indicates 100µm for H&E as well as IHC images. Figure designed by

Flavia Medeiros Savi and the author.


4.4 DISCUSSION

Large volume losses of bone substances (especially in weight bearing long bones)

resulting from multiple surgeries in fracture non-unions, revision surgeries in failed

arthroplasties or surgical oncology are a major challenge in clinical practice. They

are associated with a significant risk of an unfavourable clinical outcome (multiple

surgical interventions, high costs, and high burden of disease) often including failure

to restore full function of the affected limb. Despite multiple innovations over the

last decades current treatment options have significant limitations and there is a

strong clinical demand for novel treatment alternatives, including (off-the-shelf)

bone tissue engineering applications. Nevertheless, preclinical large animal models

of large volume long bone segmental defects (> 3cm in length) to investigate the

regenerative capacity of novel bone tissue engineering strategies under clinically

relevant conditions are rare. We herein present a newly established preclinical ovine

animal model for the treatment of large volume (6cm length) segmental tibial

defects. In addition to characterizing the mechanical properties of the scaffolds as

well as the entire construct in vitro and a detailed description of the surgical

procedure, we also present short-term (3 months) and long-term (12 months) results

of a Pilot Study using porous mPCL-TCP-scaffolds combined with a reduced dose of

2mg rhBMP-7 and PRP to regenerate this challenging defect in vivo.

4.4.1 Animal model and mechanical conditions

Based on our expertise gained in the 3cm tibial defect model, we have successfully

modified the surgical technique to create a segmental mid-diaphyseal tibial defect of

twice the length. By applying a thicker DC-plate and increasing the number of

screws used for fixation, we were able to stabilize the longer defect site with a

unicortical plate-osteosynthesis as well.

When newly establishing a bone defect animal model, it is essential to characterize

the mechanical conditions of the scaffolds or matrices used as well as of the entire

construct applied (in this case a 6cm tibial defect stabilized with a DCP and filled

with a press-fitted porous mPCL-TCP scaffold). These factors are not only important

when interpreting results of the specific study, but also to enable comparisons

between different animal models as mechanical conditions are known to substantially



influence bone healing. mPCL-TCP scaffolds with different dimensions and minor

changes (in terms of filament thickness and distance between filaments) compared to

our 3cm defect model were used in this study as outlined above. We also anticipated

increased hypoxia in the defect site due to the increased defect size and therefore

punched nine 2mm-holes aligned along the longitudinal axis of the scaffold. These

holes were placed in proximity to the neurovascular bundle upon implantation, to

allow ingrowth of larger vessels into the increased scaffold volume. Given these

modifications, scaffold mechanical properties as well as the total construct

mechanical properties have been characterized in vitro [399]. Results showed small

but statistically significant changes in scaffold stiffness properties through addition

of 9 linear holes and for testing under conditions similar to the physiological

environment after scaffold implantation. Mechanical construct testing showed

interfragmentary movement (IFM) due to functional loading to be within the range of

0.01-1 mm (0.2-2 % strain) and therefore comparable to mechanical conditions

measured in the 3 cm defect model [75]. However, the exact relationship between

mechanical conditions and bone healing remains to be elucidated and further studies

are necessary to reach consensus on this topic. A recent publication by our study

group reviewed previous methods and shared results of recent work of our group

toward developing and implementing a comprehensive biomechanical monitoring

system to study bone regeneration in preclinical tissue engineering studies [403].

Despite the relatively low results of approx. 0.01-1mm IFM from in vitro construct

testing [399], some clinical signs of increased movement in the defect site were

found in the long term observation group (group II) of this study: Implant failure

with breakage of the DCP in the defect middle occurred in one sheep where bony

non-union persisted at the end of the study. This is indicative of significant strains on

the DCP due to IFM over time with potential progressive material weakening and the

occurrence of peak loads exceeding maximum load-bearing capacity of the plate in

non-healed defects. Breakage of a single screw was detected in another sheep at three

months as well as some extent of plate bending in another sheep at three months,

furthermore indicating movement in the segmental bone gap. However, since defects

in both sheep experienced early bony bridging and IFM therefore likely decreased as

the material stiffness within the defect increased no further progressive implant

failure or total loss of internal stabilization occurred in these two sheep. All of the


above mentioned observations indicate that the mechanical environment in vivo as

well as changes over time due to healing progression need to be further analysed and

characterized. This could for example be realized by application of implantable

sensor systems (AO Fracture Monitor, AO Research Institute Davos, Davos,

Switzerland) that besides healing progression measurement in the defect site in vivo

can also provide animal activity data such as number of loading cycles within certain

time intervals and intensity ranges [403, 404]. We have recently started to investigate

the potential use of such an implantable AO Fracture Monitor (adapted for the DCP

fixation used by our research group) combined with externally fitted sheep activity

monitoring harnesses in our 6cm ovine animal model and preliminary data is

published in [403]. In future studies, maximum sheep body weight should be well

controlled and sheep activities causing peak loads should be avoided if possible.

Should implant failure become a relevant issue in the future potential modifications

of the internal fixation (e.g. addition of a second plate or changing to an

intramedullary nail as fixation device) will have to be re-evaluated.

4.4.2 Evaluation of the regenerative potential of mPCL-TCP scaffolds combined

with 2mg rhBMP7 and PRP in the novel 6cm tibial segmental defect model

In this pilot study, we have also evaluated the efficacy of a tissue engineering

construct (TEC) consisting of a mPCL-TCP-scaffold combined with a reduced dose

of 2m rhBMP-7 and PRP to regenerate this challenging tibial segmental defect. Our

results showed significant new bone formation in the defect volume and early

bridging of the complete defect length in the majority of animals at three months

after surgery, but no functional restoration of the mechanical properties. Due to the

early time point of analysis only three month after surgery and given the large

volume defect, it was to be expected that biomechnical properties of the operated

tibiae would not be significantly restored, yet. Previous studies using the 3cm tibial

defect had already revealed significantly lower values for the operated tibiae at 3

months even when applying rhBMP-7 [75], so sheep included in this study were not

expected to perform equally to the non-operated legs at this time point. But it is

noticeable that despite this relatively short healing period, values of 4-17% of max.

Torque and 5-25% of torsional moment of the non-operated tibia had already been

reached. However, these promising results at three month post implantation of the



TEC did not translate into further improved biomechanical strength and bone

volumes at twelve months after surgery. Whereas overall mechanical strength

(torsional moment) and torsional stiffness after 12 months were significantly higher

when defects were augmented with the mPCL-TCP scaffold containing rhBMP-7 in

our 3cm defect model and comparable to the unoperated control tibia [75], we found

significantly lower biomechanical properties of the operated tibiae persist after

twelve months in this study.

Furthermore, despite that total bone volumes in the defect site were increased at 12

months compared to the three month time-point, results failed to reach statistical

significance (only bone volumes in the scaffolds inner duct were significantly higher,

but not overall total bone volumes). The axial bone volume distribution pattern found

in this study was similar to results obtained in our 3cm defect model (59): In all

treatment groups a non-significant trend towards higher bone volume formation in

the proximal defect third was observed. This has been attributed to decreasing soft

tissue coverage and vascularization from proximal towards distal defect regions plus

an impaired blood supply to the distal tibial end caused by the ostectomy (59). By

trend more bone seemed to be formed in the proximal and distal defect regions

compared to the defect middle (though no statistical significance was found). This

shows that analogous to the 3cm defect, bone regeneration is initiated and propagated

in proximity to the remaining host bone at the osteotomy sites proximally and

distally, subsequently advancing towards the defect middle. Interestingly, we found

the pattern of radial bone volume distribution at three months to be different to our

3cm model: For the 3cm tibial defect the amount of newly formed bone in the

periphery in both groups was comparable to within the scaffold wall and inner duct

at 3 and 12 months. A trend towards greater bone formation in the inner scaffold duct

was observed with the addition of rhBMP-7 to the scaffold. In this study, we found

significantly higher new bone volume formation in the scaffold periphery at three

months after implantation. At twelve month, a trend towards higher bone volumes in

the scaffold periphery as well as scaffold wall was visible, but did not reach

statistical significance. Bone volumes in the scaffolds inner duct were found to be

significantly higher at twelve months compared to three months after surgery,

indicating bone formation to occur from the periphery of the defect site inwards. A

potential explanation for this observation could be that the larger defect volume

causes a significant increase of hypoxia in the defect site, favouring bone


regeneration to take place in close proximity to the adjacent soft tissue (where higher

vascularity and consequently more oxygen is available) in early stages of the defect

healing. Strong VEGF-positive staining of the soft tissue inside the defect at three

months supports the assumption of largely hypoxic conditions and strong angiogenic

signalling at early healing stages. Furthermore, we found significantly increased

bone volumes in the endosteal scaffold lumen at 12 months compared to 3 months

after surgery along with widespread presence of vWF-positive blood vessels.

A large variability of the results was observed in the twelve months group with two

animals showing fairly good restoration of mechanical properties in the range of 50%

of the non-operated leg, one sheep with moderate mechanical properties and one

sheep with complete bony non-union of the defect. Given the small sample size of

n=4 in this Pilot study, the statistical analysis of this group has therefore been most

likely compromised and a larger sample number might have been more likely to

identify outliers as well as potential differences compared to group I. However, it has

to be noted that we also observed a larger variation in bone formation in the rhBMP-

7 group in the 3cm defect model before [75]. As discussed elsewhere, potential

causes may include differences in the local mechanical environment (body weight,

individual activity and limb loading patterns etc), pH, composition and size of the

defect hematoma, minor variations in surgical technique, release kinetics, and the

concentration of local connective tissue progenitor cells or degree of vascularization

[405, 406].

Furthermore, the dosage of rhBMP-7 per mm3

scaffold volume used in this study was

significantly lower than in previous studies in the 3cm defect model: While we used

3.5mg rhBMP-7 (0.442µg rhBMP-7 per mm3 scaffold volume) in an initial study

[75], we were later able to show that a reduced dosage of 1.75mg rhBMP-7 (0.221µg

rhBMP-7 per mm3 scaffold volume) combined with the mPCL-TCP scaffold also led

to equivalent results to autograft transplantation or the high BMP dosage [78]. In this

study however, we applied a further reduced dosage of 0.133µg rhBMP-7 per mm3

scaffold volume. Our results indicate that this dosage (combined with the current

model of delivery) is too low to consistently regenerate such a challenging tibial

defect. While we found a strong osteogenic response with early defect bridging in

most animals at three months, new bone formation and mechanical properties did

overall not increase statistically significant towards 12 months after surgery

(although a trend towards higher bone volumes and increased mechanical stability



was indeed observed between the groups). BMP-7 is known to act predominantly at

early stages of fracture healing [407]. Furthermore, PRP as a carrier for growth

factors releases high protein dosages at early time points [408]. Therefore, it is likely

that the rhBMP-7 applied in this study would have led to an early and strong

osteoinductive signalling with increased doses of BMP-7 present in the

scaffold/defect volume over a short period of time. This is a potential explanation for

the strong initial healing response observed at three months after surgery. After rapid

release from PRP, BMP-7 levels would have subsided quickly due to reduced

dosage, short protein half-live and diffusion away from the defect site. Apparently,

the herin applied dose of rhBMP-7 was not sufficient to consistently initiate and

propagate bone healing processes over the entire time of the study. While the

continuing presence of the mPCL-TCP scaffold led to ongoing bone regeneration and

remodelling as seen in mCT-analyses and histology/IHC-results at 12 months, the

endogenous healing capacity combined with the osteoinductive scaffold was not

sufficient to fully regenerate the large volume defect over time despite an early short-

term osteoinductive trigger via exogenously added rhBMP-7. This hypothesis is

supported by the observation that the application of a mPCL-TCP scaffold only

(without grafting material or growth factors) led to some extent of defect healing

(compared to an empty defect), but results were significantly lower at twelve months

compared to the 3.5mg rhBMP-7-group [75].

4.4.3 Future outlook and comparison of study results with currently available

literature on ovine large segmental tibial defects

The results of this study show that large volume tibial segmental defects in ovine

animal models are challenging to treat and therefore reflect the clinical situation in

humans well. In clinical practice the use of a simple autologous bone graft in defects

of more than 5cm length is not recommended (due to significant risk of graft

resorption despite good soft tissue coverage) [56] and such defects are often treated

with bone segment transport or free vascularized bone transfers [58]. In order to offer

an alternative to consistently regenerate such extensive losses of bone substance with

tissue engineering applications, we apparently need to apply more potent stimuli

based on the diamond concept [145, 146]. Potential combinations of (smart)


osteoconductive biomaterials, (osteoinductive) growth factors and osteogenic cells

(either externally added, recruited to the defect site or provided via addition of

autologous bone grafting material) combined with sufficient vascularisation and

mechanical stability need to be investigated. Potential strategies for future studies

using our 6cm tibial segmental defect model include (but are not limited to)

combinations of mPCL-TCP scaffolds with autologous bone grafts (harvested with

the Reamer Irrigator Aspirator-System or from the Iliac crest), application of

increased doses of rhBMP-7 (similar to clinical dosage) or changing mode of growth

factor delivery (e.g. more sustained release from microparticles or alginate

hydrogels), application of BMP-2 or other growth factors (including angiogenic

growth factors).

Reviewing currently available literature on segmental tibial defects larger than 4cm

in length in ovine animal models, all previous studies had methodological

shortcomings: While results of plain radiographs and histological analyses are

published consistently, studies either lacked biomechanical testing of operated legs

[393] or accurate assessment and quantification of bone volumes in the defect site by

mCT-analyses [394] or both [312-314, 395] Only one study [396] reported both

mCT- and biomechanical test results at 4 months after surgery, but long-term data

(e.g. 12 months like in our study) are missing. Additionally, long term results for

observation periods of 1 year or more are rarely available [395]. The age of sheep

used in the studies is furthermore often not reported or relatively young sheep have

been used that do no exhibit bone microstructure similar to humans. Furthermore,

mechanical properties of applied scaffolds are often not well reported and

mechanical conditions of the entire construct in vitro or in vivo (e.g. interfragmentary

movement) are mostly not characterized.

Our animal model is to the best of our knowledge the first large volume (> 4cm

length) segmental tibial defect ovine animal model in which all the above listed

shortcomings of previous studies have been addressed. We have characterised

mechanical properties of the scaffolds as well as the scaffold construct in vitro and

are currently investigating in vivo conditions. Not only have we established a highly

standardized surgical technique in aged sheep (with Harversian remodelling). We

have also used highly standardized protocols for comprehensive ex vivo specimen

analyses (including biomechanical testing, microcomputed tomography, histological



and immunohistochemical analyses) adopted from our 3cm tibial defect model

performed by a very experienced team of researchers. We furthermore included long-

term results after twelve months in our pilot study and will continue to do so in

future experiments. Limitations of our study include the relatively small sample size

(n=4 per group), since this was a Pilot-study on technical feasibility and

characterization of the novel 6cm tibial segmental defect model. Larger sample sizes

would have increased the studies power and might have indicated potential

differences between the two groups better. Statistical results therefore should be

interpreted cautiously. Furthermore, breakage of the DCP in one sheep as well as a

single screw failure in another and some degree of DCP bending over time in a third

sheep point towards increased movement at the defect site despite our biomechanical

test results of the scaffold construct showing relatively small interfragmentary

movement in vitro. Mechanical conditions in vivo need to be analysed further. Body

weight as well as peak loading patterns must be carefully controlled in future studies

using this defect size. Addition of a second plate or an alternative fixation device

(e.g. intramedullary nail as in other studies) may also be considered, should implant

failure become a relevant problem in future studies. However, with the establishment

of this novel large volume tibial defect model and application of our well

standardized study protocols (enabling comparison between different tissue

engineering applications applied), we are able to generate valuable preclinical

evidence for the treatment of extensive segmental tibial defects.

4.4.4 Conclusion

In conclusion, we have successfully established our novel 6cm-segmental tibial

defect ovine animal model with this proof-of-concept study. We were able to

characterise the mechanical conditions of the animal model as a prerequisite for

interpreting results of this pilot study and further studies. We have also found

substantial new bone formation in the defect volume when applying porous mPCL-

TCP scaffolds combined with 2mg rhBMP-7 and PRP. We were able to characterise

bone formation patterns at early (3 months) and late (12 months) time-points in this

new model and gained valuable insight from in depth histological and

immunohistochemical analyses. However, our results show that the mPCL-TCP

scaffolds combined with a reduced dosage of 2mg rhBMP (0.133µg per mm3


scaffold volume) were not able to consistently and fully regenerate such a

challenging large volume bone defect. This is accordance with other ovine animal

studies and current clinical practice. Further studies investigating the efficacy of

increased rhBMP-7 doses or other growth factors, sustained growth factor release

and/or addition of autologous bone grafts or other bone substitute materials are

necessary to determine the most effective treatments for this challenging defect in

our new animal model.

4.5 ACKNOWLEDGEMENTS

The author would like to acknowledge Dr. Stephanie Fountain for

characterising the mechanical properties of 6cm mPCL_TCP Scaffolds and

performing scaffold construct biomechanical testing as part of her PhD-work [399].

The author would like to gratefully acknowledge all members of the QUT

histology laboratory and QUT BTM group (especially Flavia Medeiros Savi, Felicity

Lawrence and A/Prof Mia Woodruff) for their great work regarding the histological

and immunohistochemical analyses performed for this study.

Chapter 5: Final Discussion 131

Chapter 5: Final Discussion

From an orthopaedic surgeon’s point of view, segmental tibial defects with a

substantial loss of bone volume are one of the most challenging bone defects

encountered in clinical practice. While modern day medicine offers a variety of bone

substitute materials, autologous bone graft is still the gold standard against which all

other treatments need to be evaluated. Despite multiple innovations over the last

decades, no other currently clinically available bone substitute material offers an

equally efficient combination of osteoconductive three-dimensional structure,

osteogenic cells and osteoinductive growth factors with favourable mechanical

properties and vascularisation. However, graft volumes are limited and there is a

significant risk of donor site morbidity such as chronic pain and dysesthesia at the

graft-harvesting site (found in 20-30% of all patients) or bone fractures. Orthopaedic

bone tissue engineering strategies have been intensely investigated over the last

decades, but clinical translation of such applications is still rarely seen compared to

the large number of scientific studies conducted in the laboratories. Therefore,

dependable preclinical evidence generated in highly standardized large animal

studies is necessary to bridge the scale-up gap between small animal models and

translation from bench to bedside. Due to its significance in orthopaedic practice, this

research focussed on the preclinical evaluation of novel bone tissue engineering-

based treatment options for tibial segmental defects as clinical target application.

The first aim was to evaluate the regenerative potential of a novel spatiotemporal

delivery system for rhBMP-2 with extended release from a functionalized alginate

hydrogel combined with a tubular PCL-scaffold. This hybrid system had been

extensively characterised and successfully applied in a rat femoral critical defect

model before. Results from the small animal model showed extensive bone

regeneration and restoration of mechanical properties equal or superior to an

autograft control group. The first challenge of this aim was the scale-up from a small

to a large animal model which was achieved by successfully adapting scaffold design

paramters, scaffold manufacturing processes, hydrogel fabrication, hydrogel

volumes, rhBMP-2 doses and surgical techniques to be applicable in the 3cm tibial

defect ovine animal model. Results of the study showed significantly increased bone

132 Chapter 5: Final Discussion

volumes and superior restoration of mechanical properties when rhBMP-2 was

applied using the spatiotemporal hybrid delivery system compared to scaffolds

and/or hydrogels alone. Furthermore, the outcome six months after surgery was

found to be significantly better than two other recently trialled tissue engineering

constructs from the same research group. Interestingly, the herein presented study

yielded results that directly paralleled previous results from the rodent animal model.

Scientific experience shows that significant findings in small animal models are often

not reproducible in large animal models due to challenges associated with the scale-

up process (such as nutrient/diffusion changes, mass and volume challenges,

alterations in limb loading patterns and total strains applied, differences in surgical

techniques as well as changes in physiology between small and large animals). This

research study was one of the exceptional cases where a tissue engineering

application developed and tested in a small animal model has proven to be as

effectively applicable in a preclinical large animal model as well. With the combined

evidence obtained from the rodent animal model and the preclinical ovine large

animal model (both of which have been extensively characterized and analysed), a

significant step towards a potential application of the hybrid delivery system in

human trials in the near future has been accomplished.

Over the past decade various tissue engineering constructs have been investigated in

QUTs 3cm tibial defect ovine animal model. However, segmental bone defects

encountered in clinical practice are often of larger nature (especially after multiple

surgical interventions in facture non-union, after failed arthroplasties or in

orthopaedic oncology). Therefore, the second aim of this research was to establish a

large volume 6cm-tibial segmental defect model based on the expertise obtained

from the well-established 3cm-ocine tibial defect model. It was successfully shown

that modifications in surgical technique and plate fixation could be made to

accommodate a 6cm tibial mid-diaphyseal defect in the ovine animal model as well.

Biomechanical properties have been characterised (in cooperation with and by

another researcher) to further define this novel ovine animal model. As a proof-of-

concept study the regenerative potential of a previously used combination of mPCL-

TCP scaffold with BMP-7 (in a reduced dosage) and PRP has then been investigated

in the newly established large volume tibial defect model. Results of this pilot study

showed formation of substantial new bone volumes in the defect site and early defect

bridging in most animals at three month after surgery. However, long term-results at

Chapter 5: Final Discussion 133

twelve months did not yield consistent defect regeneration with insufficient

restoration of biomechanical properties of the operated tibiae compared to non-

operated contralateral bone. While the majority of defects had been bridged with

increasing volumes of mineralized bone tissue present, biomechanical properties

were not sufficiently regenerated to indicate potential full functional load bearing in

a clinical setting. Limitation of this pilot study included the relatively small sample

size of n=4 in both study groups. Therefore, statistical results should be interpreted

cautiously since larger sample sizes would have increased the studies’ power.

Nevertheless, bone healing in this novel animal model in the presence of a mPCL-

TCP scaffold and with reduced doses of BMP-7 have been extensively analysed

radiologically, histologically and immunohistochemically with valuable insights for

future studies. In summary, the second aim of this research was successfully

achieved by establishing and characterising a new 6cm-tibial defect preclinical ovine

animal model. Although the hypothesis that the application of a mPCL-TCP-scaffold

with reduced dosage of 2 mg BMP-7 would be sufficient to consistently achieve

defect restoration after 12 months had to be refuted, valuable insights for future

studies using this model were gained from this proof-of concept-study. The newly

established large segmental tibial defect was found to be challenging to treat, which

reflects the situation in humans well. Tibial segmental bone defects of more than

5cm length are amongst the most challenging bone defects encountered in

orthopaedic surgery and often require complex surgical interventions with

vascularized bone grafts or bone segment transport to achieve sufficient bone

regeneration.

Both the 3cm- and 6cm-tibial segmental defect ovine animal model presented in this

thesis are based on well-established and highly reproducible techniques performed

by a very experienced team of researchers. This enables a direct comparison between

the results of different studies testing various tissue engineering applications in each

of these ovine animal models. Due to large variations between study protocols in

vivo as well as differences in ex vivo sample analyses currently available results from

other large animal models allow at best a limited comparability regarding the

regenerative potential of different tissue engineering approaches applied. This is a

key factor compromising scientific significance of results obtained in each individual

study and ultimately hampering clinical translation. Using the 3cm- and/or 6cm-tibial

134 Chapter 5: Final Discussion

segmental defect sheep models established at QUT as a testbed will enable

researchers to directly compare results between various studies, thereby creating

highly relevant preclinical evidence on the efficacy and safety of novel tissue

engineering constructs investigated.

The “holy grail” of bone tissue engineering has not been found yet and it is likely

that no such single tissue engineering application that is capable of regenerating all

bone defects despite their origin exists. Analogous to a trend towards Personalized

Medicine in current clinical practice we will rather see the development tissue

engineering strategies developed for specific clinical target applications (e.g. tibial

segmental defects), with defined implants to be used (e.g. LISS-plating or

intramedullary nails) and applicable in certain subgroups of patients (e.g. elderly

patients with osteoporotic bone structure). This research has provided valuable

preclinical evidence on the regenerative potential of bone tissue engineering

applications in two ovine large animal models with high clinical significance. While

first clinical trials might be seen in near future for the hybrid delivery growth factor

system investigated in the first study, the best treatment options for the large volume

tibial defect model newly established in the second study remain to be elucidated.

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Appendices 157

Appendices

Appendix A

Paper 1

Statement of Contribution of Co-Authors for Thesis by Published Paper

The authors listed below have certified* that: 1. they meet the criteria for authorship in that they have participated in the conception,

execution, or interpretation, of at least that part of the publication in their field of expertise;

2. they take public responsibility for their part of the publication, except for the responsible author who accepts overall responsibility for the publication;

3. there are no other authors of the publication according to these criteria;

4. potential conflicts of interest have been disclosed to (a) granting bodies, (b) the editor or publisher of journals or other publications, and (c) the head of the responsible academic unit, and

5. they agree to the use of the publication in the student’s thesis and its publication on the QUT ePrints database consistent with any limitations set by publisher requirements.

In the case of this chapter: Computer aided design of scaffolds for bone tissue engineering

J. Henkel, J. T. Schantz , D. W. Hutmacher

Osteologie 2013, Volume 22, Issue 3, pages 180-187.

Contributor Statement of contribution*

Jan Henkel

Performed literature review, wrote the manuscript

Signature

Date 08.03.2017

Jan-Thorsten. Schantz

Aided manuscript preparation

Dietmar W. Hutmacher

Aided literature review and manuscript preparation

QUT Verified Signature

158 Appendices

Principal Supervisor Confirmation

I have sighted email or other correspondence from all Co-authors confirming their

certifying authorship.

Dietmar W. Hutmacher _ 27.07.2017

Name Signature Date


Appendices 159

Appendix B

Paper 2








In the case of this chapter: Bone Regeneration based on Tissue Engineering Conceptions – A 21st Century

Perspective

J. Henkel, M. Woodruff, D.R. Epari, R. Steck, V. Glatt, I.C. Dickson, P.F. Choong, M.A.

Schuetz, D.W. Hutmacher

Bone Research (2013) 3: 216-248.


Jan Henkel

Performed literature review, wrote the manuscript, designed figures

Signature

Date 08.03.2017

Maria Woodruff Aided manuscript preparation

Devakar Epari Aided manuscript preparation

Roland Steck Aided manuscript preparation


160 Appendices

Vaida Glatt Aided manuscript preparation

Ian Dickinson Aided manuscript preparation

Peter Choong Aided manuscript preparation

Michael Schuetz Aided manuscript preparation






Dietmar W. Hutmacher _ 27.07.2017

Name Signature Date


Appendices 161

Appendix C

Paper 3








In the case of this chapter: Design and fabrication of scaffold-based tissue engineering

J. Henkel, D.W. Hutmacher

BioNanoMaterials, Volume 14, Issue 3-4, Pages 171–193, December 2013


Jan Henkel

Performed literature review, wrote the manuscript, designed figures

Signature

Date 08.03.2017






Dietmar W. Hutmacher ________________ 27.07.2017

Name Signature Date

162 Appendices

Appendix D

Paper 4








In the case of this chapter: Delayed Minimally Invasive Injection of Allogenic Bone Marrow Stromal Cell Sheets

Regenerates Large Bone Defects in an Ovine Preclinical Animal Model.

Berner A*, Henkel J*, Woodruff MA, Steck R, Nerlich M, Schuetz MA, Hutmacher DW

Stem Cells Transl Med. 2015 Apr 1. pii: sctm.2014-0244.

*(both authors contributed equally)


Jan Henkel

Conducted experiments and data analysis, wrote the manuscript, designed figures

Signature

Date 08.03.2017

Arne Berner Experimental design, conducted experiments, wrote the manuscript

Maria Woodruff Aided data analysis and manuscript preparation

Roland Steck Aided data analysis and manuscript preparation

Michael Nerlich Aided manuscript preparation

Appendices 163

Michael Schuetz Aided data analysis and manuscript preparation


Experimental design, data analysis, manuscript preparation





Name Signature Date

164 Appendices

Appendix E

Paper 5








In the case of this chapter: Scaffold-cell bone engineering in a validated preclinical animal model: precursors vs

differentiated cell source

Berner A, Henkel J, Woodruff MA, Saifzadeh S, Kirby G, Zaiss S, Gohlke J,

Reichert JC, Nerlich M, Schuetz MA, Hutmacher DW.

J Tissue Eng Regen Med. 2015 Dec 9.


Jan Henkel

Conducted data analysis, wrote the manuscript, designed figures

Signature

Date 08.03.2017

Arne Berner Experimental design, conducted experiments, wrote manuscript

Woodruff Mia Aided data analysis and manuscript preparation

Siamak Saifzadeh Conducted experiments

Giles Kirby Aided data analysis and manuscript preparation

Appendices 165

Sascha Zaiss Conducted experiments, aided manuscript preparation

Jan Gohlke Conducted experiments, aided manuscript preparation

Michael Nerlich Aided manuscript preparation

Michael Schuetz Aided manuscript preparation


Experimental design, data analysis, manuscript preparation





Name Signature Date

166 Appendices

Appendix F

Paper 6








In the case of this chapter: Monitoring Healing Progression and Characterizing the Mechanical Environment in

Preclinical Models for Bone Tissue Engineering

Fountain S, Windolf M, Henkel J, Tavakoli A, Schuetz MA, Hutmacher DW, Epari DR.

Tissue Eng Part B Rev. 2015, Dec 15.


Jan Henkel

Conducted experiments, aided manuscript preparation

Signature

Date 08.03.2017

Stephanie Fountaint Experimental design, conducted experiments, data analysis, wrote manuscript

Markus Windolf Conducted experiments, data analysis, wrote manuscript

Aramesh Tavakoli Conducted experiments, data analysis, aided manuscript preparation

Michael Schuetz Aided data analysis and manuscript preparation

Appendices 167


Experimental design, data analysis, aided manuscript preparation

Devakar Epari Experimental design, data analysis, aided manuscript preparation





Name Signature Date

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