“Explore everything, stop at nothing”

—Lara Croft

List of Papers

This thesis is based on the following papers, which are referred to in the text by their Roman numerals.

I Piskounova, S., Forsgren, J., Brohede, U., Engqvist, H. and

Strømme, M. (2009) In vitro characterization of bioactive tita-nium dioxide/hydroxyapatite surfaces functionalized with BMP-2. Journal of Biomedical Materials Research, Part B, Ap-plied Biomaterials, 91B(2):780–787.

II Piskounova, S., Rojas, R., Bergman, K. and Hilborn, J. (2009) The effect of mixing on the mechanical properties of hyaluro-nan-based injectable hydrogels. Macromolecular Materials and Engineering, in press; doi: 10.1002/mame.201100008.

III Piskounova, S., Gedda, L., Hulsart-Billström, G., Bergman, K., Hilborn, J. and Bowden, T. Characterization of recombinant human bone morphogenetic protein-2 delivery from injectable hyaluronan-based hydrogels by means of 125I-radiolabeling. Submitted manuscript.

IV Piskounova, S., Hulsart-Billström, G., Gedda, L., Bergman, K., Larsson, S., Hilborn, J. and Bowden, T. Pre-incubation of chemically crosslinked injectable hyaluronan-based hydrogels loaded with rhBMP-2 and its effect on ectopic bone formation. Submitted manuscript.

V Hulsart-Billström, G., Piskounova, S., Gedda, L., Hilborn, J., Andersson, B-M., Bergman, K., Larsson S. and Bowden, T. Improved bone formation through increased surface area of hyaluronan-based hydrogels. Manuscript.

Reprints were made with permission from the respective publishers.

The following papers are not included in this thesis:

VI Ossipov, D., Piskounova, S. and Hilborn, J. (2008) Poly(vinyl alcohol) cross-linkers for in vivo injectable hydrogels. Macro-molecules, 41(11):3971–3982.

VII Bergman, K., Engstrand, T., Hilborn, J., Ossipov, D., Piskou-nova, S. and Bowden, T. (2009) Injectable cell-free template for bone-tissue formation. Journal of Biomedical Materials Re-search, Part A, 91A(4):1111–1118.

VIII Ossipov, D., Piskounova, S., Varghese, O., and Hilborn, J. (2010) Functionalization of hyaluronic acid with chemoselec-tive groups via a disulfide-based protection strategy for in situ formation of mechanically stable hydrogels. Biomacromole-cules, 11(9):2247–2254.

IX Forsgren, J., Brohede, U., Piskounova, S., Mihranyan, A., Lars-son, S., Strømme, M. and Engqvist, H. (2011) In vivo evalua-tion of functionalized biomimetic hydroxyapatite for local deli-very of active agents. Journal of Biomaterials and Nanobio-technology, 2:150–155.

X Koens, M.J.W., Geutjes, P.J., Hoogenkamp, H.R., Ainödhofer,

H., Wismans, R.G., Tiemessen, D.M., Faraj, K.A., Piskounova, S., Mindemark, J., Hilborn, J., Gupta, B., Feitz, W.F.J., Daa-men, W.F., Saxena, A.K., Oosterwijk, E., van Kuppevelt, T.H. Construction and in vivo vascularization of large-diameter tubu-lar scaffolds for gastrointestinal regenerative medicine. Manu-script.

My contributions to the papers in this thesis

I I contributed to the design of the study, performed all of the cell experiments, participated in the discussion of the results and wrote the main part of the manuscript.

II I contributed to the design of the study, performed all of the ex-periments, participated in the discussion of the results and wrote the main part of the manuscript.

III I contributed to the design of the study, performed all of the ex-

periments and wrote the manuscript.

IV I contributed to the design of the study, performed the rheologi-cal characterization and in vitro release experiments, partici-pated in the discussion of the results and wrote the main part of the manuscript.

V I contributed to the design of the study, performed the in vitro

release experiments and parts of histological analysis, partici-pated in the discussion of the results and wrote a major part of the manuscript.

Contents

1 Introduction ......................................................................................... 13 1.1 Osseointegration and bone regeneration ......................................... 13 1.2 Principles of bone formation and remodeling ................................. 14 1.3 Current strategies and materials ...................................................... 16

1.3.1 Hydroxyapatite....................................................................... 16 1.3.2 Bone morphogenetic proteins ................................................ 17

1.3.2.1 Discovery of bone morphogenetic proteins ................ 17 1.3.2.2 Bone morphogenetic protein-2 ................................... 18 1.3.2.3 Carrier systems ........................................................... 19

1.3.3 Hyaluronan hydrogels ............................................................ 20

2 Results and discussion ......................................................................... 22 2.1 Functional implant coatings ............................................................ 22

2.1.1 Coating preparation ................................................................ 22 2.1.2 Cell viability, diffentiation and morphology ......................... 22

2.2 Hydrogels ........................................................................................ 25 2.2.1 Experimental design .............................................................. 25 2.2.2 Hydrogel preparation ............................................................. 28

2.2.2.1 The effect of mixing ................................................... 28 2.2.2.2 Hydrogel swelling in different buffers ....................... 32 2.2.2.3 The effect of addition of rhBMP-2 and HAP ............. 33

2.2.3 Delivery of rhBMP-2 from hydrogels .................................... 34 2.2.3.1 Optimization of protein handling and detection ......... 34 2.2.3.2 In vitro release and ALP activity ................................ 35 2.2.3.3 In vivo release ............................................................. 38 2.2.3.4 Bone and cartilage formation ..................................... 41

3 Concluding remarks ............................................................................. 51

4 Future prospective ............................................................................... 53

5 Acknowledgements.............................................................................. 56

6 Svensk sammanfattning ....................................................................... 60

7 References ........................................................................................... 63

Abbreviations

ALP Alkaline phosphatase bFGF Basic fibroblast growth factor BMC Bone mineral content BMD Bone mineral density BMP-2 Bone morphogenetic protein-2 BMV Bone mineral volume CHO Chinese hamster ovary ECM Extracellular matrix ELISA Enzyme-linked immunosorbent assay GAG Glycosaminoglycan GMSA Gel mobility shift assay HA Hyaluronic acid HAA Aldehyde derivative of HA HAP Hydroxyapatite MSC Mesenchymal stem cell MTT Thiazolyl blue tetrazolium bromide PBS Phosphate-buffered saline PVA Poly(vinyl alcohol) PVAH Hydrazide derivative of PVA pQCT Peripheral quantitative computer tomography

Scope of the thesis

One of the greatest challenges in modern medicine is tissue failure or loss caused by injury or disease, which results in reduced quality of life for pa-tients who are afflicted by such conditions. The lack of available donor tis-sue has inspired the initiation of a new interdisciplinary field, which is a combination of materials science, chemistry and biology, referred to as “tis-sue engineering”. Within this field medical doctors and scientists work to-gether to repair or replace damaged or unhealthy tissue, as well as to im-prove the growth and function of new healthy tissue.

Bone, unlike many other tissues in the body, has the exceptional ability to heal itself. However, in the cases when the tissue’s own regenerative capaci-ty is not sufficient, bone healing and bone formation can be enhanced via various tissue engineering approaches. By using specific biomaterials in combination with bioactive molecules, it is possible to stimulate and enhance the body’s own regenerative pathways in a manner similar to the processes of natural bone formation and remodeling.

In this thesis, two pathological conditions are addressed. The approaches implemented for treating these conditions, as well as the way the different projects are interconnected, is described in Scheme 1.1. The first condition, addressed herein, is poor bone/implant integration, caused by the poor bone-bonding nature of the commonly used titanium implant surface. In Paper I strategies are presented for preparing multilayered functional coatings of titanium implants for stimulating initial osseointegration via osteoconductive hydroxyapatite (HAP) and osteoinductive bone morphogenetic protein-2 (BMP-2), as well as improving long-term bone/implant-bonding via a layer of crystalline TiO2. The characterization of the in vitro cellular response on the aforementioned surfaces shows that these coatings have great potential and that biomolecules can be loaded into the HAP layer by a simple soaking method. Furthermore a possible interaction between HAP and BMP-2 is observed.

The second condition, addressed in this thesis, is severe bone damage. The strategy for stimulating bone formation at ectopic sites involves deliver-ing BMP-2 via an injectable chemically crosslinked hyaluronan hydrogel. First of all, a reproducible hydrogel preparation technique, based on the mix-ing of two polymer components, is presented and the appropriate release medium and setting time are selected for these hydrogels in Paper II. Se-condly, a reliable technique for BMP-2 handling and detection are discussed

in Paper III. In the same body of work the BMP-2 in vitro and in vivo release kinetics are characterized and correlated to protein activity and ectopic bone formation in rats. Furthermore, the in vitro and in vivo effects of pre-incubation of premixed hydrogels prior to injection are studied in Paper IV. Finally, the implication of variation of surface area on the material perfor-mance is explored by comparing solid hydrogels to crushed hydrogels in Paper V.

This thesis highlights the importance of detailed characterization of the individual components for the design of reliable medical devices for bone regenerative applications.

Scheme 1.1. A flowchart over the correlations between Papers I–V.

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1 Introduction

1.1 Osseointegration and bone regeneration This thesis addresses two major pathological conditions involving bone tis-sue. The first is implant loosening due to poor integration between the im-plant and the adjacent bone, i.e. osseointegration (Figure 1.1a). The second is bone regeneration in patients with severe bone injury or bone loss, as illu-strated in Figure 1.1b. These two conditions present major challenges in both dentistry and orthopedics, and are therefore more closely described herein.

Titanium is frequently used to manufacture load-bearing implants for clinical applications, due to its bioinert nature and high mechanical strength. However, since bone does not directly bond to titanium, implant loosening may occasionally occur and cause supplementary bone loss under mechani-cal load.1 Since the average human life expectancy is constantly increasing, patients receiving e.g. artificial hip joints and knee prostheses require an improved and prolonged implant performance, in order to have a chance of an active lifestyle.2 The current average lifetime of such load-bearing im-plants is only 10–15 years and revision surgery is often required earlier, due to infection and poor implant fixation.3, 4 In general, these procedures are more complicated than the primary surgery, requiring the removal of the old implant and the surrounded fibrous tissue.5, 6

Figure 1.1. Two major challenges in orthopedics and dentistry: a) poor osseointe-gration and b) regeneration of severe bone defects.

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Unlike many other tissues, bone possesses an excellent self-healing ability. In most cases, with appropriate reduction and fixation of bone fractures, new bone formation can be observed just a few weeks after an injury. Provided that there are no complications, total bone union and complete functional recovery is reached after a few months. However, in severe pathological situations, such as complicated fractures, birth defects, bone tumor extrac-tion, joint replacement, spinal fusion etc., the damaged bone does not form or regenerate spontaneously.7 Hence, in search for alternative treatments, the field of tissue engineering and regenerative medicine have been employed. This interdisciplinary field was founded in the 1990’s in an attempt to repair or replace damaged or unhealthy tissue, as well as improve the growth and function of new healthy tissue, by combining the basic principles of mate-rials science, chemistry and biology.8 One of the strategies of bone tissue engineering and bone regeneration is to mimic the natural processes of bone formation and bone remodeling, which are presented more thoroughly in the following section.

1.2 Principles of bone formation and remodeling Bone is a unique type of connective tissue which is both strong and light. In addition to its supportive and protective function, bone acts as a reservoir for inorganic ions, it plays e.g. an important role in the homeostasis of calcium in the body. The extracellular matrix (ECM) of this tissue is made up of both a non-mineralized phase (osteoid), containing type I collagen and glycosa-minoglycans (GAGs), and a mineralized phase, comprising calcium phos-phate salts.9

Figure 1.2. a) Different types of bone tissue. b) Intramembranous and endochondral ossification.

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Most bones in the body are comprised of two morphologically and function-ally different forms of bone (Figure 1.2a). The outer layer consists of com-pact bone, also referred to as cortical bone. Compact bone is made up of densely packed collagen fibrils, which form concentric rings (lamellae). The lamellae are organized in perpendicular planes, such as in plywood, giving compact bone its rigid properties suitable for mechanical support. The inner layer is composed of a loosely organized and porous trabecular bone, also referred to as cancellous or spongy bone. Trabecular bone, which surrounds the bone marrow, has a more metabolic function.9

During embryonic development, bone formation is achieved via direct or indirect ossification (Figure 1.2b). Flat bones of the skull, the mandible, maxilla, and clavicles are formed via intramembranous ossification, during which mesenchymal stem cells (MSCs) differentiate directly into osteogenic cells.9, 10 On the contrary, long bones and other load-bearing bones form via endochondral ossification, during which MSCs first differentiate into carti-lage, which is subsequently replaced by weak woven bone with randomly organized collagen fibers. Woven bone is gradually replaced by mature and rigid lamellar bone, consisting of highly organized concentric layers of inor-ganic and organic ECM.9

Figure 1.3. Bone remodeling is dynamic process based on a balance between bone resorption and bone formation.

In adults, bone continues to undergo constant remodeling in response to me-chanical load, changes in local calcium concentrations and a wide range of paracrine and endocrine factors. This dynamic process regulates the balance between bone formation and resorption, which occurs in both compact and trabecular bone, and is governed by bone-producing cells (osteoblasts) and bone-resorbing cells (osteoclasts), as illustrated in Figure 1.3. The remode-ling starts by osteoblasts triggering osteoclasts to break down bone matrix. The activated osteoclasts form a ruffled border in the area of bone resorption and release organic acids and lysosomal enzymes to break down the inorgan-ic and organic bone components, respectively. As a result of this process,

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calcium is released into the bloodstream. Over time, osteoblasts replace os-teoclasts and begin to lay down new lamellar bone on top of the old bone. So-called cement lines mark the borders between the old and the new bone matrix.9, 11

1.3 Current strategies and materials Modern medicine provides several options for patients with severe bone loss, of which autologous bone grafting is considered to be the gold standard. In general, during this treatment, bone from the iliac crest of the patient’s own pelvis is transplanted to the site of the defect. Despite the prevalence of this technique it has severe drawbacks, such as donor site morbidity, additional pain and the risk of infection. Furthermore, the autologous grafts can often be insufficient due to the limited supply of suitable donor tissue from one individual, anatomical and structural problems and high resorption levels. Another option is surgical transplantation of bone tissue from other donors (allografts). However, this requires the patients to be treated with immuno-suppressive drugs during the remainder of their lives. Treatment with allo-grafts also introduces the risk of disease transmission.12, 13

The aforementioned limitations of the existing treatments motivated the search for alternative materials, which would enhance bone formation via one or several of the three basic mechanisms: osteoconduction, osteoin-duction and osteogenesis. During osteoconduction bone formation is trig-gered from existing bone by the introduction of a scaffold with a three-dimensional structure similar to that of bone tissue. Osteoinduction implies that new bone is formed by endogenous or transplanted osteoprogenitor cells in response to biomolecules, e.g. growth factors. Osteogenesis can be in-itiated by manipulating the natural process of bone remodeling, e.g. by inhi-biting bone resorption.14

1.3.1 Hydroxyapatite One of the most frequently used osteoconductive materials in orthopedic and dental surgery is hydroxyapatite (HAP). This bioactive, biocompatible and biodegradable calcium phosphate Ca10(PO4)6(OH)2 resembles the mineral phase of bone and has excellent mechanical properties. Therefore, in order to mimic bone mineral composition, porosity and strength, HAP has often been employed to produce fillers or bioactive implant coatings on dental and or-thopedic implants.15, 16

Furthermore, it has been shown that HAP may have specific affinity to-wards certain osteogenic17 and anti-osteoresorptive molecules,18 and that particles, as well as coatings, may be used to create reservoirs for growth factors,19 antibiotics and osteoclast-inhibiting drugs.18, 20 Such coatings are

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particularly interesting, as they could allow relatively simple patient-specific treatment. One such strategy for using HAP together with biological mole-cules as bioactive coatings of titanium implants is discussed in Paper I.

1.3.2 Bone morphogenetic proteins 1.3.2.1 Discovery of bone morphogenetic proteins Bone morphogenetic proteins (BMPs) were discovered in 1965 by an ortho-pedic surgeon, Dr. Marshall R. Urist, who showed that extracts of deminera-lized bone could induce ectopic bone and cartilage formation when it was implanted in non-skeletal tissue in rats.21, 22 This discovery commenced a new research era, the main goal of which was to identify, purify and employ new growth factors to regenerate damaged or missing tissue. Some of the most frequently used growth factors in modern regenerative medicine are listed in Table 1.1.

Table 1.1. Most commonly used growth factors. Modified from Chen (2003).23

Growth factor Abbreviation Physiological function

Basic fibroblast growth factor bFGF/FGF-2 Proliferation of fibroblasts and initia-tion of angiogenesis

Bone morphogenetic protein BMP-2 BMP-7 (OP-1)

Differentiation and migration of bone-forming cells

Epidermal growth factor EGF Proliferation of epithelial, mesen-chymal, and fibroblast cells

Platelet-derived growth factor PDGF-AA PDGF-AB PDGF-BB

Proliferation and chemoattractant agent for smooth muscle cells; extra-cellular matrix synthesis and deposi-tion

Transforming growth factor- TGF- Migration and proliferation of kerati-nocytes; extracellular matrix synthe-sis and deposition

Transforming growth factor- TGF- Proliferation and differentiation of bone-forming cells; chemoattractant for fibroblasts

Vascular endothelial growth factor VEGF Migration, proliferation, and survival of endothelial cells

The isolation and characterization of BMPs proved to be laborious, since 1 kg of bone only contained 1–2 µg of these proteins.24 Through introduction of dissociative extraction25 and advances in molecular biotechnology, it was finally possible to determine the amino acid sequences of each individual purified protein.26 Several BMPs were successfully isolated and characte-rized by Wang and coworkers in 1988.27 A year later the same scientists patented the first recombinant human BMPs (rhBMPs) produced by trans-fecting the full length cDNA via mammalian expression vectors into Chinese hamster ovary (CHO) cells.27

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At present, approximately 20 different BMPs have been identified. As other members of the transforming growth factor beta superfamily, BMPs act as homo- or heterodimers, which are held together by disulfide bonds. It has been shown that BMPs are not limited to bone formation, but play an impor-tant mechanistic role in other tissues and organs as well.28, 29 However, in dental and orthopedic applications, where bone induction is desired, BMP-2 has received immense attention, due to its unique ability to promote osteo-genic differentiation.10, 30

1.3.2.2 Bone morphogenetic protein-2 The molecular action of BMP-2 (Figure 1.4) is triggered when the protein binds extracellularly to two transmembrane receptors, BMPR-I and BMPR-II, and a cascade of intracellular events via SMAD proteins is initiated, re-sulting in activation of numerous osteogenic genes, such as collagen type I, osteocalcin, osteopontin and alkaline phosphatase (ALP).31 BMP-2 can in-duce bone formation via both intramembranous and endochondral ossifica-tion pathways. In rodents, low doses result in the formation of small amounts of cartilage, which is later replaced by bone, while large doses lead to direct bone formation.10

Figure 1.4. BMP-2 signaling pathway where BMP-2 binds to transmembrane recep-tors BMPR-I and BMPR-II, initiating a chain of intercellular events, resulting in activation of osteogenic genes.

The recombinant BMP-2 is excreted from CHO cells as a 30 kDa dimeric glycoprotein.26 The highly basic isoelectric point (pI = 8.5) gives rhBMP-2 a positive net charge at physiological pH. The highly hydrophobic surface of the protein and its unique sensitivity towards changes in ionic strength makes rhBMP-2 virtually insoluble in water and PBS.32 Therefore, in order to prevent protein aggregation and precipitation, a concentrated HCl solution

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or an acidic formulation buffer* are recommended for all use of the pro-tein.32–34

1.3.2.3 Carrier systems The use of rhBMP-2 in clinical applications proved to be complicated due to the protein’s short biological half-life of only a few minutes, its localized action and its rapid clearance. To solve this issue a carrier system, which allows for sustained and local rhBMP-2 delivery, is required.12 Various ma-terials and systems (Table 1.2) have been proposed, with the common goal to retain rhBMP-2 in its active form at the site of implantation long enough for the MCSs to migrate in and initiate differentiation into osteoblasts.35, 36

Table 1.2. Examples of materials and carriers used for BMP-2 delivery.

Materials Examples

Ceramics Hydroxyapatite37, 38 Bioglass39 Tricalciumphosphate (TCP)38

Natural biopolymers Collagen40 Fibrin41, 42 Gelatin43–45 Alginate46 Chitosan47 Hyaluronan48–50

Synthetic polymers Poly(lactic acid) (PLA)51 Poly(glycolic acid) (PGA)52 Poly(lactic-co-glycolic acid) (PLGA)53

Composites PLA–hyaluronan54 Chitosan–gelatin55 Poly(caprolactone) (PCL) –collagen–TCP56

In recent years several medical devices in conjunction with BMPs have emerged. In 2002 a product based on the delivery of rhBMP-2 absorbed in a type I collagen sponge, was approved for anterior lumbar interbody spinal fusion in the US by the Food and Drug Administration and marketed as IN-FUSE® Bone Graft kit (Medtronic).57 That same year, this collagen/rhBMP-2 formulation was approved in Europe by the European Medicines Agency for the treatment of open tibial shaft fractures and is now sold under the product name InductOs® (Pfizer, former Wyeth).40 Several difficulties have been reported with both products. First of all, due to the poor affinity of the colla-gen towards the rhBMP-2 molecule, this carrier shows limited protein reten-tion. Thus, large doses on a milligram scale are required for clinical effect, making these treatments expensive.40, 58 Secondly, the bovine origin of the collagen introduces additional risks of immunological reaction and inflam-

*5 mM L-glutamic acid, 2.5 % L-glycine, 0.5 % sucrose, 5 mM NaCl and 0.01 % polysor-bate 80, pH 4.5.

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mation.59, 60 Lastly, these products need to be implanted surgically, which can cause additional tissue damage and fibrosis. Therefore, an alternative non-toxic, non-immunogenic carrier material, which can be delivered via a non-invasive technique, e.g. via injection, and which has a specific affinity for rhBMP-2, is required.

1.3.3 Hyaluronan hydrogels Hyaluronan (HA) was first isolated by Meyer and Palmer in 1934 from the vitreous of bovine eyes. This negatively charged, non-sulfated, linear GAG is abundant in nature, found in the connective tissues of all mammals. The HA molecules from different species have the same primary structure com-prised of a repeating disaccharide unit, as shown in Figure 1.5, and only vary with regards to the number of repeating units, making HA highly bio-compatible.61

Figure 1.5. The chemical structure of HA, consisting of the repeating disaccharide unit of 4)- -D-GlcAp-(1 3)- -D-GlcNAcp-(1 .

In addition to its structural role, HA is involved in several biological processes,61 such as cell differentiation,62 angiogenesis,63–66 morphogenesis,67 and wound healing.68, 69 These features, together with HA’s biocompatibility and its excellent biophysical properties, e.g. viscoelasticity and pseudoplas-ticity*, make this biopolymer appropriate for a number of medical applica-tions.70 For example, native HA is commonly used during viscosupplementa-tion and viscosurgery,71–74 and as an adjuvant for ophthalmologic† drug deli-very.75 HA’s ability to inhibit osteoclast differentiation,76 in addition to its down-regulation of BMP-2 antagonists,77 makes it an interesting material to use for bone-regenerative purposes.

Today, large quantities of highly pure HA are available in a wide range of molecular weights and at a relatively low cost. HA is mostly extracted from rooster combs or produced through microbial fermentation. However, native HA, even of high molecular weight, is a soluble substance, which lacks me-chanical stability and is rapidly degraded and cleared in vivo.61 To overcome this and to obtain materials with tailored properties, various approaches have been employed to chemically modify HA.78–80 Modified HA has been used to

*Pseudoplasticity means that the material becomes less viscous in response to shear force. †Ophthalmologic drugs are drugs used to treat eye disease.

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prepare scaffolds for tissue engineering applications,81–86 for soft and hard tissue regeneration,87–89 and for controlled drug and growth factor deli-very.90–92

A group of hydrogels, based on chemical crosslinking between an alde-hyde derivative of HA (HAA) and a hydrazide derivative of poly(vinyl alco-hol) (PVAH), as shown in Figure 1.6, was proposed as a delivery vehicle for rhBMP-2 by Bergman et al. in 2009, showing promising results in an ectopic bone model in rats.93 The same material, loaded with HAP in addition to rhBMP-2, was successfully used to heal critical-size defects in rats and pigs.94, 95 Despite the great potential of this hydrogel system in vivo, there remained many unanswered questions, in addition to the need for reproduci-ble methods for handling and characterization of these scaffolds. Some of these questions and requirements were addressed in Papers II–V.

Figure 1.6. An aldehyde derivative of HA (HAA) and a hydrazide derivative of poly(vinyl alcohol) (PVAH) react rapidly and selectively to form hydrazone cross-links.

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2 Results and discussion

2.1 Functional implant coatings 2.1.1 Coating preparation Paper I presents an in vitro characterization of multilayered functional im-plant coatings, designed to stimulate immediate bone formation as well as long-term osseointegration (Figure 2.1). For this, circular plates of commer-cially pure grade 5 titanium were coated with a gradient layer of bioactive crystalline titanium dioxide (TiO2) via physical vapor deposition. On top of the TiO2 a porous HAP layer was deposited with phosphate-buffered saline (PBS) as the ion source. Finally the HAP was loaded with rhBMP-2 via a simple soaking step. The rationale behind this design was to employ the osteoinductive properties of rhBMP-2 and the osteoconductive nature of HAP to promote initial bone formation. Furthermore, it was hypothesized, that once the biodegradable HAP is resorbed, the underlying layer of bioac-tive crystalline TiO2 would be more favorable for osseointegration than the native titanium oxide (TiO) that forms spontaneously on titanium surfaces upon exposure to oxygen.

Figure 2.1. Multilayered functional coating proposed for improved initial and long-term osseointegration.

2.1.2 Cell viability, diffentiation and morphology The in vitro response, in terms of cell viability, differentiation and morphol-ogy, on different bioactive coatings was studied in Paper I. The W20-17 clone of MSCs was selected for this study, since this particular cell line has been shown to exhibit a direct correlation between cell differentiation and the amount of added rhBMP-2. The cell differentiation was measured by

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quantifying the increase in ALP expression by a colorimetric assay.96 Since BMP-2 does not affect the proliferation of W20-17 cells, cell viability was only studied on surfaces without the growth factor, i.e. native TiO, crystal-line TiO2 and HAP. Cell differentiation and morphology, however, were studied on HAP, loaded with rhBMP-2, in addition to the aforementioned surfaces.

Figure 2.2. a) Cell viability was measured by an MTT assay and presented as the number of cells detected on native TiO, crystalline TiO2 and HAP surfaces after 2 d of cell culture. b) Cell differentiation was determined by an ALP assay and pre-sented as the absorbance at 405 nm after 1 day and 2 d of culture on native TiO, crystalline TiO2, HAP and HAP loaded with rBMP-2. Values are presented as means ± standard deviation of the mean for n = 3.

Results showed that crystalline TiO2 was more favorable than the native TiO surface in terms of cell viability (Figure 2.2a). Morphologically, the W20-17 cells formed more spider-like networks on crystalline TiO2 than on native TiO, suggesting better cell attachment (Figure 2.3). Finally, cell differentia-tion (Figure 2.2b), was also higher on crystalline TiO2. The fact that the cells favored crystalline TiO2 over native TiO could be due the increased wettability and surface charge on TiO2 surfaces. Wettability has been pre-viously demonstrated to have a positive effect on cells.97 The electrostatic interaction between the OH-terminated and deprotonated crystalline TiO2

surface and the positive calcium ions from the cell culture medium could also increase cell attachment.

The HAP surface demonstrated an even higher cell viability than the crys-talline TiO2, suggesting that the cells preferred the tree-dimensional surface of the HAP crystals. Despite the increased number of cells and their exten-sive cell spreading, HAP alone did not stimulate significant cell differentia-tion.

The cell differentiation on HAP, loaded with BMP-2, was found to be significantly higher than on all the other surfaces. The fact that rhBMP-2 could be loaded into the HAP layer via a simple soaking procedure indicates a possible interaction between the inorganic material and the growth factor.

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Molecular modeling has previously shown that HAP crystals may have spe-cific binding sites for rhBMP-2.17

In addition to the aforementioned data, the preliminary results of a quartz crystal microbalance study, where the binding of rhBMP-2 to HAP-coated crystals was investigated, indicated that after ~40 min, the growth factor bound irreversibly to the surface (data not shown).

The phenomenon of a possible HAP/rhBMP-2 interaction is highly im-portant for designing a carrier for rhBMP-2. First of all, the short loading time and simple procedure indicate that HAP surfaces can be used for on-demand incorporation of active molecules. The orthopedic or dental surgeon could on the day of the surgery choose to include rhBMP-2 loading to pro-mote faster initial osseointegration in a specific patient. Furthermore, it could be possible that HAP is able to bind endogenous BMPs, thus stimulat-ing osteoinduction as well as osteoconduction. The latter information could be of value for the design of biomaterials for treating severe bone defects.

Figure 2.3. Scanning electron microscopy images of the morphology of W20-17 cells, cultured for 2 d on a) native TiO, b) crystalline TiO2, c) HAP and d) HAP loaded with BMP-2.

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2.2 Hydrogels 2.2.1 Experimental design For treating severe bone damage and bone loss, chemically crosslinked hy-drogels were used to deliver osteoinductive and osteoconductive components to the site of injury. When designing such a system, it is important to explore and standardize everything from material preparation to the biological out-put. Furthermore, just as during the design of the multifunctional implant surface, it is important to take into account the capabilities and the require-ments of the end-users, which, in the case of this work, are the orthopedic and dental surgeons. The material should be easy to use and any possible deviation from the standard procedure that may change the performance of the material should be well-characterized and underlined to minimize the risk of device failure and additional patient suffering.

With this in mind, in Papers II–V, the implications of variation of several parameters on the performance of the hydrogels as a delivery vehicle for rhBMP-2 were investigated. The hydrogels used in this work included the aforementioned HAA/PVAH system93 and the AuxigelTM (TERMIRA, Stockholm, Sweden). Apart from a slight variation in polymer synthesis and polymer molecular weight, the underlying mechanism in both systems is the same. This mechanism involves a rapid crosslink formation, which occurs when the two polymers are mixed. Osteoinductive biomolecules, such as rhBMP-2, and osteoconductive fillers, such as HAP particles, can be incor-porated into the hydrogel by addition to the hydrogel components prior to mixing.93–95, 98

In Paper II, in order to have better control over the aforementioned deli-very system, the preparation method for hydrogels without any additives was first standardized by optimizing the mixing procedure and determining the ideal setting time. Hydrogel samples were prepared by the mixing of the polymer solutions back and forth for 5, 20 and 50 mixing cycles, where one mixing cycle was defined as the passing of all the material from one syringe to the other. The resulting constructs were referred to as 5×, 20× and 50× hydrogels and the total duration time of each preparation technique was 2.5, 10 and 25 s, respectively. Each polymer premix was injected into custom-made aluminum forms, directly after mixing, and rheological measurements were performed, to determine the difference in mechanical properties caused by the different number of mixing cycles (Figure 2.4). By keeping the hy-drogels in the form it was possible to maintain constant sample shape and temperature (37 °C), reduce evaporation, and allow for simple sample trans-fer to an incubator between measurements, without damaging the hydrogel. Most importantly, it was possible to perform repeated rheological measure-ments of the same hydrogel sample during extended periods, thus saving time and material.

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The hydrogel stability during long-term in vitro experiments, e.g. during the study of rhBMP-2 release kinetics, was of concern. Even well-mixed and fully crosslinked hydrogels could undergo excessive swelling in certain buf-fers, leading to hydrogel rupture and dissolution. Moreover, such swelling would not correspond to the in vivo scenario, where significantly less water would be present and the swelling would most likely be prevented by the pressure of the surrounding tissue. Hence, once the mixing and the setting time of the hydrogels had been selected, the effect of different release media on hydrogel swelling was explored.

Figure 2.4. A schematic representation of the hydrogel preparation by mixing of the two components (left) and experimental set-up for rheological measurements (right).

During an in vitro release kinetics study, limited amounts of rhBMP-2 could be detected by the biochemical technique used by Bergman et al.93 Further-more, protein loss due to aggregation and adhesion to surfaces during this experiment was a feared to be a problem. Therefore, in Paper III, reliable methods for rhBMP-2 handling and detection were developed, including the selection of optimal sample tubes with minimal rhBMP-2 affinity.

By employing these attained methods for hydrogel preparation and rhBMP-2 handling and detection, a thorough characterization of the growth factor release from 0.2 mL hydrogels, loaded with 30 µg of rhBMP-2, was conducted in Paper III. The in vitro release and retention of rhBMP-2 was followed for 28 d by radiolabeling a fraction of the growth factor with 125I and measuring the counts in the release medium and in the hydrogel, respec-tively. Furthermore, the ALP expression of W20-17 cells, cultured with the hydrogel release medium, was studied to determine the bioactivity of the released protein. The radiolabeling technique was also used to follow the rhBMP-2 retention in vivo. For this, the material was injected in the thigh muscle of male Sprague–Dawley rats. At different time points, the animals were sacrificed, and the amount of retained rhBMP-2, as well as the biodi-

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stribution of 125I in different organs, was measured and correlated to the ec-topic bone formation, which was analyzed by radiographic analysis and pe-ripheral quantitative computer tomography (pQCT).

During the aforementioned study a question arose regarding what would happen should the hydrogel be injected minutes, hours, or even days after premixing. This question was highly relevant from the clinical perspective, since the surgical environment is quite unpredictable, and hence a premixed system could be more convenient. Therefore, the possibility of hydrogel preparation prior to initiation of surgery, or even delivery of the material as a premix, was of interest. To investigate this, a subcutaneous ectopic bone model was chosen, as described in Paper IV, where each 0.2 mL hydrogel was loaded with 50 mg of HAP, in addition to 30 µg of rhBMP-2. This for-mulation and animal model were selected to make it easier to retrieve the hydrogel grafts after several weeks in vivo.

First of all, the effects of rhBMP-2 and HAP on the curing time were ex-plored, by using the rheological technique described in Paper II. Based on the results, three pre-incubation times were selected (1 min, 5 h and 3 d) in order to produce slightly, moderately and fully crosslinked hydrogels. The in vitro release of the growth factor from these constructs was followed by radioactive labeling as described in Paper III and the in vivo tissue formation was characterized by radiographic analysis, pQCT, histology and immuno-histochemistry. This study revealed that the “crushing” of the hydrogel net-work, which occurred upon injection of the moderately and fully crosslinked hydrogels resulted in scaffolds with a significantly larger surface area than injection of slightly crosslinked hydrogels. Since several advantages of scaf-folds in particulate form were know from literature, including porosity, pore size, vessel infiltration, etc.,99, 100 the implications of increasing the surface area were further explored in Paper V, by a comparing solid hydrogels and crushed hydrogels in vitro and in vivo. The scaffolds used in this study were manufactured according to the procedure described in Figure 2.5. The in vitro release of rhBMP-2 from each hydrogel type was studied using radi-olabeling and the biological effect of the increased surface area was explored by subcutaneously implanting solid hydrogels and crushed hydrogels on the back of rats. The grafts were retrieved after four weeks and assessed radio-graphically, as well as by pQCT and histology.

The findings of Papers II–V, regarding the optimal hydrogel preparation, rhBMP-2 handling, detection and delivery, as well as the resulting tissue formation, are discussed in more detail in the following sections of this the-sis.

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Figure 2.5. Preparation of hydrogels with different surface area in Paper V. a) Component A (HAA and HAP) and component B (PVAH, HAP, rhBMP-2) were loaded into two luer-lock syringes. b) The syringes were connected by an adapter and the components were premixed to form a hydrogel. The hydrogel premix was used to manufacture solid and crushed hydrogels. c–e) For solid hydrogels the hy-drogel premix was injected into cylindrical forms directly after mixing, and was allowed to cure for 24 h. f–h) Crushed hydrogels were prepared by first leaving the material to cure for 24 h inside the one of the syringes that had been used for mix-ing. The next day the hydrogels were pushed through the tip of the syringe twice.

2.2.2 Hydrogel preparation 2.2.2.1 The effect of mixing Several factors may affect the structural and mechanical properties of poly-meric hydrogels, including chain entanglement and crosslinking density, polymer molecular weight and concentration, crosslinker length, polymer mixing, curing time, osmotic pressure, additional fillers, biological mole-cules, etc.

A common approach for preparing injectable chemically crosslinked hy-drogels is by mixing two or more reactive components. Such hydrogels can form in situ, if the mixing takes place during injection, or they can be pre-formed by mixing prior to injecting. Currently a wide range of different sy-ringe systems is available on the market. To meet the numerous require-ments of material developers, several companies have developed syringes

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with various types of mixing tips (both static and dynamic) in order to op-timize mixing, including dual- or even triple-barreled syringes with differ-ent-sized cartridges and mixing tips.

The advantages of the back-and-forth method, used to prepare hydrogels in Papers II–V were many. First of all, unlike many commercially available cartridges, the syringes could be filled at different occasions and with the volume of choice. Secondly, this technique allowed “on-demand” addition of sensitive biomolecules. This was particularly valuable, since some sensitive biomolecules need to be kept in their natural buffer for as long as possible for best effect.33 The biomolecules could be added at the last minute and in small quantities, due to the minimal volume loss. Furthermore, the injections of the resulting hydrogels could be made at different occasions, which could be important for the tuning of the hydrogel mechanical properties, as de-scribed in Paper IV. Finally, it was possible to control the injected volume. Thus, one hydrogel premix could be used for several injections, which proved to be very convenient during e.g. in vitro experiments, as it saved both time and material.

The main weakness of the back-and-forth method was the variations in material performance, associated with a non-standardized mixing procedure. Therefore, in Paper II, the effect of mixing on the structural and mechanical properties was investigated, in order to develop a mixing procedure which would result in stable, well-crosslinked hydrogels.

By performing oscillatory stress sweeps on the 5×, 20× and 50× samples at different time points it was possible to follow the curing profile of each hydrogel type. The results showed that at the time of the first measurement, which was performed approximately 5 min after mixing, the average storage modulus G' (Figure 2.6a) was larger than the average loss modulus G'' (Figure 2.6b), indicating that the polymer premix had already transformed from a viscous liquid to an elastic solid through the process of gelation. The gel point, when G' = G'', was approximated to be reached after 20–60 s. This meant that for the 50× hydrogels, which took 25 s to prepare, the mixing continued past the gel point.

The stiffness of the hydrogels increased over time, suggesting ongoing curing. During the first hour G' increased significantly faster for the 20× and the 50× hydrogels in comparison to the 5× hydrogels. Furthermore, G'' stabi-lized faster for the 50× hydrogels than for the other hydrogel types. This indicated that better mixing resulted in an improved distribution of function-al groups, leading to better crosslinking. On the other hand, the 50× samples, which appeared to reach equilibrium after 3 h, were less stiff than the 20× samples after 24 h, with G' almost as low as for the 5× samples. This signi-fied that by mixing beyond the gel point, some of the formed bonds were broken, resulting in weaker hydrogels.

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Figure 2.6. Results of the oscillatory stress sweeps of hydrogels: a) G' as a function of time t; b) G'' as a function of time.

In addition, frequency sweeps were performed on 5×, 20× and 50× samples in order to describe the stability and the microstructure of the resulting hy-drogels, e.g. covalent bonds that are broken and reformed during longer measurements, and physical entanglements that are created. G', as a function of frequency, exhibited a plateau in the range of 0.1 1 Hz for all hydrogel types, suggesting stability over this frequency range (Figure 2.7a).

By plotting the logarithms of G' and the angular frequency in this fre-quency range, it was possible to further describe the kind of interactions formed by each mixing procedure (Figure 2.7a). The slope of this graph corresponded to the degree of frequency dependence, n, where n = 0 for a completely covalently crosslinked hydrogel, and n > 0 for a physical gel.101 For 5×, 20× and 50× samples the n values were low, implying that all three techniques led to the formation of almost fully covalently crosslinked hydro-gels. The 20× preparation gave rise to hydrogels with the highest mechanical stability and the most covalent crosslinks (n 0).

Figure 2.7. Results of the frequency sweeps: a) G' as a function of frequency and log G' as a function of log ; b) tan as a function of frequency.

The loss tangent (tan ) is the measure of energy loss per frequency sweep cycle and is a way to view the shear modulus as a complex quantity com-prised of a real (G') and an imaginary part (G''). Therefore, tan < 1, as was true for all three hydrogels (Figure 2.7b), additionally confirmed that the

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resulting materials were more elastic than viscous. Furthermore, from the minimum of tan , and in the linear viscoelastic region, the G'p could be identified, which was the peak value of G'. This point, where G' was almost constant, indicated that the modulus was dictated by purely elastic elements. Although this was an oversimplified representation of the crosslinked net-work studied, G'p could be used to calculate Mc, which is the molecular weight of chain segments between crosslinks, from Mc = c RT/G'p, where c is the polymer concentration (1 % w/w), is the density of water at 37 °C (992 kg/m3), R is the gas constant (8.314 J · K 1 · mol 1) and T is the temper-ature (310 K). The resulting Mc values of 5×, 20× and 50× samples were 30, 21 and 23 kg/mol respectively, implying that the 5× and 50× hydrogels were less densely crosslinked than the 20× hydrogels.

A summary of the mechanistic implications of the different mixing pro-cedures is presented schematically in Figure 2.8. Based on rheological mea-surements of the hydrogels, prepared by different numbers of mixing cycles, it could be concluded that the 5× samples were the most loosely crosslinked, probably due to the fact that the small number of mixing cycles prevented the reactive functional groups from finding each other. In the case of the 50× samples, although these hydrogels had reached equilibrium the fastest, they proved to be weaker than the 20× hydrogels, suggesting that excessive mixing beyond the gel point resulted in the irreversible breaking of cross-links and other bonds.

Figure 2.8. A summary of the implications of the different mixing procedures. In-sufficient mixing gives rise to not fully crosslinked hydrogels, while excessive mix-ing leads to breaking of the formed crosslinks and other chemical bonds.

The events taking place during excessive mixing beyond the gel point could be described as three interdependent stages. The hydrogel system had a con-

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siderable number of functional groups (F) per polymer chain, thus the con-version (P) at the gel point would be very low (P = 2/F), as suggested by Carothers.102 Hence, in the first stage of mixing beyond the gel point, even if some of the formed bonds were partially ruptured, the crosslinking was con-tinued by the residual functionalities, which through continuous mixing were able find each other and link the fragments of the broken hydrogel. The ca-pability of this re-linking thereby became dependent on the mixing, which decreased the diffusion distances between free functional groups, allowing them to react. In the second stage, at higher conversion and a decreased number of available reactive groups, continuous mixing led to an increased number of broken covalent bonds. Finally, in the third stage the hydrogel structure was dominated by clusters of broken hydrogel with little possibili-ties of re-linking due to a low content of remaining unreacted functional groups.

Although the ideal number of mixing cycles remained to be determined, based on these results, it could be concluded that stable hydrogels could be formed by mixing between 20 and 50 times. Therefore, in Papers III–V, the hydrogels were prepared by mixing 30 times. Furthermore, from the curing profiles of the hydrogels, it was concluded that 3 h of setting time, prior to the addition of swelling or release medium, was sufficient for minimizing the risk of hydrogel dissolution due to insufficient crosslinking time.

2.2.2.2 Hydrogel swelling in different buffers In Paper II, the swelling of the hydrogels in PBS and in serum-containing cell culture medium was compared. These two buffers were regarded as the most likely to be used during further in vitro work, since they resemble the physiological environment.

The results showed that 20× hydrogels swelled significantly more in cell culture medium than in PBS (Figure 2.9a). This could be due to the pres-ence of additional molecules and proteins in the cell culture medium. These molecules could disrupt inter- and intramolecular entanglement by compet-ing for interaction with polymer functional groups, via hydrogen bonding and through electrostatic interactions. They could also increase the chances of hydrolysis of bonds, by aminolysis. Moreover, the interaction of proteins with the hydrogel matrix could drive more water molecules into the system over time, which would result in an apparent increased swelling.

Comparison of the swelling in PBS of hydrogels prepared by the different number of mixing cycles showed, in agreement with the results of the rheo-logical data, that an insufficient number of mixing cycles promoted forma-tion of less densely crosslinked constructs, and that excessive mixing beyond the gel point resulted in damaged hydrogels that were more susceptible to swelling.

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Figure 2.9. Swelling kinetics of the hydrogels prepared by different numbers of mixing cycles. a) The swelling ratio S of the 5×, 20× and 50× samples in PBS and of 20× in cell culture medium as a function of time; b) Schott’s second order linear regression of swelling. The values are presented as means ± standard deviation of the mean for n = 3.

By looking at the linear regression of the data, a linear relationship was found between the reciprocal of the average swelling rate t/S and the swel-ling time t, indicating that the swelling process followed Schott’s second order swelling kinetics.103 As shown in Figure 2.9b, the swelling occurred in two phases; during the first week, the samples underwent swelling to equili-brium, while during the remaining three weeks of measurements the hydro-gels underwent some hydrolysis/dissolution of chains and physical dissocia-tion of entanglements.

None of the hydrogels dissolved during the length of the experiment. However, it was decided that in order to avoid extended swelling, the release experiments should be performed in PBS. It was also decided that, since the release experiments were to be performed at 37 °C to resemble body temper-ature, the fresh buffer should be brought to this temperature prior to replac-ing the old buffer.

2.2.2.3 The effect of addition of rhBMP-2 and HAP The method for performing rheological measurements using the custom-made form, described in Section 2.2.2.1, was used in Paper IV to character-ize the mechanical properties of rhBMP-2-loaded hydrogels, with and with-out HAP. The results, presented in Figure 2.10, showed that the addition of HAP increased the material stiffness significantly. It was also observed that hydrogels containing HAP showed higher mass loss due to evaporation (up to 13 % after 24 h, vs. 5 % in hydrogels without HAP). The G' values for both hydrogels continued to increase after 24 h of measurements. However, this time-dependent increase in modulus was associated with the drying of the hydrogels during longer measurements.

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Figure 2.10. Results of the oscillatory stress sweeps of rhBMP-2-loaded hydrogels, prepared by 30 mixing cycles with and without HAP. Data is presented as the aver-age G' ± standard deviation of the mean for n = 3 vs. time and shows an increase of material stiffness upon addition of HAP.

2.2.3 Delivery of rhBMP-2 from hydrogels 2.2.3.1 Optimization of protein handling and detection The in vitro release of rhBMP-2 from the HAA/PVAH hydrogels was pre-viously described by Bergman et al.93 During this experiment, 0.2 mL hy-drogels, loaded with 30 µg of rhBMP-2, were placed in 5 mL tubes and were allowed to cure for 1 h at room temperature. Thereafter, each hydrogel was covered with 4 mL of PBS, and incubated at 37 °C. The release was fol-lowed for 28 d, where at each time point the old medium was collected and replaced with fresh PBS. The collected samples were stored at 20 °C. After 28 d the samples were thawed, diluted to appropriate concentrations in PBS and the amount of released rhBMP-2 was quantified using an Enzyme-Linked Immunosorbent Assay (ELISA). This biochemical assay is based on protein recognition by specific antibodies. The results showed that roughly 12 % of the loaded rhBMP-2 was released after 28 d. The remaining 88 % was assumed to remain inside the hydrogel.93 However, upon digestion of the hydrogels with hyaluronidase the remaining growth factor could not be detected (unpublished data).

The sensitivity of rhBMP-2 is a key factor, which should be taken into account when designing a release or bioactivity experiment, as well as dur-ing general handling and storage of this growth factor. It is known from lite-rature that rhBMP-2 easily aggregates and undergoes conformational changes at physiological conditions.32, 33 Consequently, it is possible that the release of rhBMP-2 into PBS, the freezing/thawing, and the diluting, had led to the fact that a fraction of rhBMP-2 was no longer detectable by anti-bodies or that some of the protein had simply adhered to the surface of sam-ple tubes and pipette tips.

A more reliable quantification technique was desired in order to properly characterize the release of rhBMP-2 from these hydrogels. Therefore, in

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Papers III–V, radiolabeling by 125I was selected, since this method made it possible to quickly quantify the growth factor at every step, without the need for dilutions or the use of expensive and time-consuming assays.

Furthermore, prior to performing the release experiments, the radioactive labeling was used in Paper III to select the appropriate tubes and conditions for the handling and storage of rhBMP-2. The results showed that Protein LoBind tubes exhibited the lowest protein affinity in comparison to glass vials and siliconized or regular Eppendorf tubes. It was also concluded that storage of rhBMP-2 release samples at reduced temperatures should be avoided.

2.2.3.2 In vitro release and ALP activity Once the optimal techniques for hydrogel preparation and rhBMP-2 handling and detection were selected, the in vitro release of the growth factor could be examined by radioactive labeling, as described in Paper III. The use of ra-dioactive labeling, rather than ELISA, permitted to quantify at each time point the amount of protein that was released (Figure 2.11a), as well as the amount that was retained inside the hydrogel (Figure 2.11b). The good cor-relation between release and retention profiles suggested minimal protein loss during the release experiment, probably due to the use of LoBind tubes for sample handling. A biphasic release was observed, with a fast rhBMP-2 discharge during the first week, followed by a slower release for the remain-ing three weeks of the experiment. After 28 d ~80 % of rhBMP-2 had been released, which is significantly more than was earlier detected by ELISA.93 The initial fast release took place at about the same time as the swelling of the hydrogels in PBS.

Figure 2.11. The in vitro release of rhBMP-2 from the hydrogels: a) cumulative release and b) retention of 125I-labeled rhBMP-2. These profiles of the labeled pro-tein were assumed to correspond to those of the unlabeled rhBMP-2.

In order to better understand this biphasic release profile, the dimensions of rhBMP-2 and the average pore size of the hydrogel were more closely ex-amined. Based on the rheological measurements in Paper II, the average poor size (mesh size) of the hydrogels was calculated in Paper III to be

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around 52 nm, and from the crystal structure of rhBMP-2, the dimensions of the rhBMP-2 dimer were 7 × 3 × 3 nm.104 However, it has been stated in literature that this growth factor aggregates to 100 nm particles even in for-mulation buffer and that the size of these aggregates at pH 6.5 increases to up to 2 µm.33

Therefore, the protein released during the initial fast phase was probably composed of small non-aggregated rhBMP-2 complexes that gradually dif-fused out of the hydrogel during the swelling process, during which the mesh size increased. Larger aggregates were retained within the polymer network due to their size.

Naturally, there remained a question whether the released growth factor was biologically active. In order to check this, an ALP assay was performed on the release samples to determine if the released rhBMP-2 could still trig-ger differentiation of W20-17 cells into MSCs. The results showed an initial increase in ALP activity till day 14, followed by a decrease till day 28 as shown in Figure 2.12.

Figure 2.12. ALP activity of the rhBMP-2 release samples. Error bars represent means ± standard error of the mean for n = 3. Statistical significance: * p < 0.5.

It was, however, not possible to directly compare the release experiment and ALP experiments since they were performed under slightly different condi-tions. During the release experiment the old release medium was replaced with fresh PBS at each time point, thus possibly driving the rhBMP-2 out of the hydrogel. On the other hand, for the ALP experiment fresh hydrogels were prepared at each time point, covered with PBS and left till the last day, when all of the release medium was collected and used to culture cells. The different procedure was chosen for the ALP experiment in order to avoid storage, pooling and different dilutions of all the release samples. The alter-native to perform a new ALP experiment for each time point was tested (data not shown), but large variations in cell behavior, due to differences in pas-sage and confluence made the results difficult to interpret.

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Nevertheless, if for the sake of argument it would be assumed that the re-lease profile during the ALP experiment was the same as during the release experiment, then more information could be extracted regarding the stability of rhBMP-2. For example, the fast increase in ALP expression correlated to the increase in rhBMP-2 concentration due to the fast release during that time. Furthermore, a large decrease in ALP activity, observed between day 14 and day 28, could be due to the fact that very little fresh rhBMP-2 was being released from the hydrogel during that time, while the already released protein was gradually inactivated in the physiologic environment of the release buffer and prolonged storage at 37 °C.

In Papers IV–V hydrogels loaded with HAP, in addition to rhBMP-2, were investigated. The goal of the study in Paper IV was to investigate the implication of a pre-incubation of the hydrogel premix prior to injection in an attempt to demonstrate different scenarios which may occur in the operat-ing room. For this slightly, moderately and fully crosslinked hydrogels were manufactured by pre-incubating the hydrogel premix for 1 min, 5 h and 3 d, and compared.

Interestingly, there was no significant difference in rhBMP-2 retention between the different hydrogel groups (Figure 2.13a). However, morpholog-ical differences were observed, where injection after 1 min gave rise to a compact scaffold, and the 5 h and the 3 d hydrogels resembled thin entan-gled threads. After a few days these string-like structures appeared to partial-ly fuse together.

A possible explanation for this phenomenon would be that injection of the hydrogels, after they had reached the gel point, as was done for the 5 h and 3 d samples, could result in “crushing” of the three-dimensional network. This crushing, in conformity with the events observed after excessive mix-ing, would result in breaking of the formed crosslinks, but also in the expo-sure of unreacted functional groups. Thus, such scaffolds, with initially high surface area, would with time form chemical bonds between the crushed fragments, resembling the compact structure of the 1 min hydrogel. Conse-quently, the diffusion distance of rhBMP-2, at least in vitro, became similar in all three pre-incubation groups.

A similar trend of the hydrogel fusing and similar release was observed for the solid and the crushed hydrogels, compared in Paper V. Despite the difference in initial surface area, the two types of constructs had virtually identical in vitro release kinetics (Figure 2.13b).

Interestingly, the release of rhBMP-2 from the hydrogels with HAP in Papers IV and V was significantly slower than from the hydrogels without HAP in Paper III (Figure 2.13c). The possible interaction between rhBMP-2 and HAP, which had been suggested to take place for the titanium coatings in Paper I, could be the reason for this additional retention in hydrogels with HAP. Furthermore, it is possible that the presence of HAP changed several

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physical parameters of the hydrogel, e.g. porosity, pore size and water reten-tion, thus increasing the rhBMP-2 retention.

Figure 2.13. a) In vitro retention of rhBMP-2 in slightly, moderately and fully cros-slinked hydrogels with HAP (1 min, 5 h and 3 d), described in Paper IV. b) In vitro retention of rhBMP-2 in solid and in crushed hydrogels, described in Paper V. c) A comparison of the average in vitro retention profiles of rhBMP-2 from hydrogels without HAP, used in Paper III and the hydrogels with HAP, used in Papers IV and V.

2.2.3.3 In vivo release Through 125I-labeling of rhBMP-2 it was possible to follow the release of this growth factor from the hydrogels in vivo. For this, in Paper III, hydro-gels, containing a mixture of 125I-labeled and non-labeled rhBMP-2, were injected into the thigh muscle of rats directly after mixing of the hydrogel components, and the retention of the growth factor was followed for 28 d, as shown in Figure 2.14a. The biphasic release profile, observed in vivo, had a significantly faster initial phase (day 0–3), with only ~45 % rhBMP-2 re-tained after 2 d, as opposed to the ~65 % that remained after this time in vitro.

There are several explanations for this. Firstly, in the body the scaffolds would be subjected to enzymatic action, e.g. via hyaluronidase, which would lead to accelerated hydrogel degradation. Secondly, during the study of the

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in vitro release the hydrogels had been allowed to cure for 3 h at 37 °C in-side tubes, prior to the addition of the release medium. This incubation step was added in order to minimize the risk of rapid influx of water into the hy-drogels, which could result in excessive swelling due to insufficient cross-linking. Such pre-incubation was not possible in vivo, since injection after 3 h would risk breaking the formed crosslinks. Furthermore, due to mechani-cal forces the same degree of swelling was not expected in vivo. Therefore, for the in vivo release experiment the material was injected directly after the mixing of the hydrogel components.

Figure 2.14. a) In vivo release kinetics of rhBMP-2 from hydrogels without HAP indicated that 45 % of the loaded rhBMP-2 was retained after 2 d and after 28 d only ~3 % remained in the hydrogels. b) Biodistribution 125I in different organs. The values are presented as means ± standard error of the mean for n = 6 and n = 3, re-spectively.

During the second phase (day 3–28) rhBMP-2 was released at a slower, more controlled rate, indicating a possible chemical interaction between the hydrogel and rhBMP-2, in addition to the size-dependent retention, which had been observed in vitro. It is possible that the formation of bone inside the hydrogel at after ~8 d, discussed in more detail in Section 2.2.3.4, slowed down the release of the growth factor, by creating a dense barrier around the graft. Furthermore, as it was mentioned in the introduction to this thesis, the pI of 8.5 means that rhBMP-2 should have a positive net charge at pH 7.4. Therefore, it could bind to the negatively charged HAA component via electrostatic interactions. Furthermore, the binding could be additionally improved via reversible Schiff base formation between the unreacted alde-hyde groups of the HAA and the amines of the protein.

Since most of the rhBMP-2 was released during the first few days in vivo, after four weeks only approximately 3 % of the initial growth factor re-mained inside the thigh muscle of the animals. Therefore, if slower release is to be desired, the system needs to be optimized. However, given the availa-bility of a reliable detection method via radiolabeling, the release profiles from different hydrogel formulations could be easily compared both in vitro

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and in vivo, with the intention to develop a more efficient delivery system for rhBMP-2 and minimize the necessary therapeutic dose. It should also be said, that the complexity of the in vivo environment makes it difficult to pre-dict the behavior of a hydrogel system only based on in vitro experiments, as shown in Papers IV and V. The information acquired in vitro and in vivo should be considered complementary rather than comparable. Therefore, both pieces of the puzzle are necessary to fully understand what happens to rhBMP-2 inside and outside the hydrogel.

In addition to detection of 125I-labeled rhBMP-2 retained in the thigh muscle, the radioactive counts were measured in the liver, kidneys and thy-roid, to determine the biodistribution of 125I in the body. Two days after in-jection, 2 and 5 % of the initial injected activity were detected in the liver and thyroid, respectively (Figure 2.14b). The somewhat higher accumula-tion of 125I in the thyroid was an indication of metabolized protein, where radioiodine is released. No significant increase in 125I levels was detected in the kidneys.

Furthermore, the amount of rhBMP-2 that was lost during each injection was of interest. To have better control over the injected volume, the syringes were intentionally loaded with enough material to prepare 0.3 mL of hydro-gel, of which 0.2 mL was injected, and 0.1 mL remained in the connector, needle and syringes. The radioactivity at t = 0 in the syringes, needle and connector was compared to the values detected at this point in the thigh. The results, presented in Table 2.1, were normalized to the total loaded radioac-tivity in the 0.3 mL of hydrogel.

Table 2.1. The average amount of 125I-labeled rhBMP-2 that was injected in rats and that remained in the syringes, connector and needle.

Localization Fraction of total loaded

hydrogel volume (0.3 mL)

Fraction of total loaded amount of 125I-labeled

rhBMP-2 (45 µg)

Injected 66 % 54 %HAA syringe + connector 14 %PVAH and rhBMP-2 syringe 33 % 25 %Needle 7 %

Surprisingly, although ~33 % (0.1 mL) of the total hydrogel volume had not been injected, ~46 % of the total loaded 125I-labeled rhBMP-2 and, hence, of the total rhBMP-2 (~21 µg), remained inside the syringes, connector and needle. This meant that 13 % (~7 µg) of the growth factor was unintentional-ly lost along the way. A large fraction of the total loaded radioactivity (25 %), which corresponded to ~11 µg of the total rhBMP-2, was found in the syringe, in which the growth factor had been loaded together with the PVAH component. A possible explanation to this was that rhBMP-2 had adhered to the syringe when the solution, containing the growth factor, was

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stored inside these for several hours prior to mixing and the subsequent for-mation of hydrogels. Therefore, the binding of rhBMP-2 to the syringes should be characterized, as was done for the sample tubes in Paper III. For now, with this data in mind, it could be concluded that bone formation in-duced by these hydrogels, and described in the following section, was in-duced by ~23 µg of rhBMP-2 and not 30 µg as was intended.

2.2.3.4 Bone and cartilage formation The ability of the rhBMP-2 hydrogels with and without HAP to promote ectopic bone formation was investigated in Papers III–V. As mentioned in the previous section, in Paper III an intramuscular ectopic rat model was chosen, in order to compare the results to the ectopic bone formation ob-served by Bergman et al. in 2009.93 As shown in the radiographic images in Figure 2.15a newly formed bone was observed next to the femur already after 8 d. The presence of mineralization at this point was additionally con-firmed by pQCT (Figure 2.15b–e). Although the exact chain of biological events remains to be investigated, it could be hypothesized that the initial fast release of rhBMP-2, observed during the first week (Figure 2.14a), was sufficient to recruit MSCs to the site of injection and initiate their differen-tiation into osteoblasts. Furthermore, as mentioned in the previous section, the formation of bone could have slowed down the release of rhBMP-2 from the scaffold, by creating a dense barrier, which could be observed on the radiographs, particularly after 8 and 14 d (Figure 2.15a). After 28 d the newly formed bone had become more well-defined and compact, signifying increased bone density. Bone mineral density (BMD), bone mineral content (BMC) and bone mineral volume (BMV) were calculated from pQCT data by complementary software (Figure 2.15c–e).

The data showed that BMV and BMC indeed increased with time, sug-gesting continued osteogenic action, probably stimulated by the retained growth factor at the site of injection. However, although BMD increased significantly between 14 d and 28 d, a decrease in BMV was observed at the latter time point. A possible explanation for this is that the 3 % of the loaded rhBMP-2 remaining inside the hydrogel was no longer sufficient to stimulate continuous bone formation. According to literature, in the absence of growth factors, non-functional bone tissue is gradually resorbed.105

The results of the in vivo experiment in Paper IV, the experimental set-up of which is described in more detail in Section 2.2.1, showed that pre-incubation of the hydrogel premix prior to injection had a large effect on the hydrogel degradation profile and the morphology of the formed tissue. The radiographic images of the grafts, explanted after five weeks (Figure 2.16), show compact round structures for the 5 h and 3 d groups and significantly less organized and partially crushed morphology for the 1 min. This sug-gested that the pre-incubation step prior to injection gave rise to hydrogels

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that were more resistant to mechanical load from the surrounding tissue, than when the material was injected directly.

Figure 2.15. Intramuscular ectopic bone formation induced by hydrogels loaded with only rhBMP-2. a) Radiographic images after 8, 14 and 28 d, where newly formed bone is indicated by arrows. b) Representative pQCT sections from of bone formation at each time point. c) BMD, d) BMC and e) BMV as functions of time, determined from the obtained pQCT data for each time point. Values are presented as means ± standard error of the mean for n = 6. Statistical significance: *** p < 0.001; ** p < 0.01. The data indicates the presence of ectopic bone forma-tion after 8 d, with increasing BMC values at the later time points. The decrease in BMV after 28 d suggests initiation of bone resorption.

The results of the pQCT analysis (Figure 2.16) showed a decrease in BMD for the 3 d group compared to the 1 min and 5 h groups. The BMC was sig-nificantly higher in the 5 h group than the 1 min group. Both the 5 h and 3 d groups had significantly higher BMV than the 1 min group. These results suggested that slightly crosslinked hydrogels induced less bone formation, possibly due to the fact that these hydrogels had poor mechanical stability.

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This mechanical stability increased with pre-incubation time, resulting in an increasing bone volume for the moderately and fully crosslinked hydrogels. However, the fully crosslinked hydrogels were less osteogenic than the slightly and moderately crosslinked hydrogels, possibly due to a decrease in rhBMP-2 activity. Therefore, although bone covered a large volume in the 3 d grafts, it was not as mature and hence not as dense as in the 5 h grafts.

Figure 2.16. (Top) Radiographic images of the 1 min, 5 h and 3 d grafts 5 weeks after subcutaneous injection of slightly, moderately and fully crosslinked hydrogels on the back of rats. (Bottom) Results of pQCT including BMD, BMC and BMV values determined for each test group. Values are presented as means ± standard deviation of the mean for n = 5. Statistical significance: * p < 0.05; ** p < 0.01; *** p < 0.0001. One graft from the 3 d group had been unintentionally injected into an intradermal area, rather than into a subcutaneous location, and was therefore excluded from this study. The results of the radiographic analysis showed that the 5 h and 3 d grafts had a more organized and compact morphology than 1 min grafts. The results of the pQCT analysis showed that the 1 min pre-incubation resulted in the least BMV, but high BMD, the 5 h pre-incubation gave rise to a higher BMV, with comparable BMD; and the 3 d pre-incubation led to formation of the highest BMV, but the lowest BMD. A possible explantation for this difference is that longer pre-incubation resulted in an increase in mechanical strength of the hydrogels and a decrease in rhBMP-2 stability.

From the first look at this data, it could be speculated that 3 d pre-incubation was a bit too long and a shorter time, e.g. 1 d, would be more ideal. Howev-er, before drawing any definite conclusions, the quality and the morphology of the formed tissue in response to the slightly, moderately and fully cross-linked hydrogels were examined in more detail by histological analysis and immunohistochemistry.

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The results of histological analysis, presented in Figure 2.17, revealed that although new cancellous bone was present in all groups, a significant mor-phological diversity was observed in the formed tissue as well as in the hy-drogel degradation profile.

Figure 2.17. Histological evaluation of the grafts after five weeks in vivo, following subcutaneous injection of slightly, moderately and fully crosslinked hydrogels, loaded with both rhBMP-2 and HAP. Representative histological sections stained with Masson’s Trichrome, from the a) 1 min, e) 5 h and i) 3 d groups shown in full view (scale bar: 1 mm), as well as at higher magnification of the b–d) 1 min, f–h) 5 h and j–l) 3 d (scale bar: 100 µm). The images show residues of hydrogel (G) and hydroxyapatite (HAP), as well as formation of non-mineralized bone (B), minera-lized bone (MB) and bone marrow (BM). The bone marrow was often filled with red cells (RC), which were presumably hematopoietic cells and MSCs. All grafts were surrounded by a layer of fibrotic tissue (F). In the 3 d group a presence of cartilage was observed, including chondrocytes (C) and chondroblasts (CB).

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The information regarding the hydrogel stability and organization, obtained from the radiographs of the 1 min grafts, was accordant with the fragile and unorganized structures observed during the histological analysis of this group, presented in Figure 2.17a–d. The 1 min grafts were generally smaller than the 5 h and 3 h ones, which also correlated to the BMV values obtained from pQCT (Figure 2.16). The 1 min grafts were surrounded by a thin peri-pheral layer of mineralized bone, inside which “islets” of hydrogel residues, as well as of mineralized and non-mineralized cancellous bone matrix, could be distinguished. In between these islets, adipocytes of the bone marrow (stained white) were observed, as well as other cells, which were presumably hematopoietic and mesenchymal progenitor cells. These progenitor cells were stained red by the hematoxylin of the Masson’s Trichrome kit and will therefore be referred to as “red cells” herein. Blue spider-web-like structures were also found inside these grafts and were interpreted to be HAP residues.

Despite the extrusion through the needle during injection, the 5 h and 3 d hydrogels had formed compact spherical scaffolds, presumably due to the fusion of the crushed fragments, as discussed in Section 2.2.3.2. The organi-zation of these grafts resembled that of an onion. In the 5 h grafts, the outer layer of bone was thicker than for the other groups, but less mineralized than in the 3 d group. Inside were a layer of HAP and a layer of red cells, as shown in Figure 2.17e–h. In the center of the onion-like structure of the 5 h graft was a large hydrogel residue, which in contrast to the “islets”, observed for the 1 min grafts, appeared to be intact, and was in most cases completely cell-free.

In the very middle of the 3 d grafts was a cell-free hydrogel residue, as in the 5 h group. However, more adipocytes and fewer red cells were observed in the 3 d grafts. The bone shell, surrounding the 3 d grafts, was thinner than that of the 5 h grafts (Figure 2.17i–l). Inside this shell were a layer of adipo-cytes and a layer of HAP filled with red cells. Surprisingly, further inside the graft the layer of HAP gradually transformed into a layer of chondrocytes and a layer of chondroblasts (Figure 2.17j). Immunohistochemical staining for aggrecan and collagen type II confirmed the presence of cartilage in this area (Figure 2.18). This information suggested that in the 3 d grafts bone formation took place via the endochondral ossification pathway, rather than through direct ossification. This is consistent with literature, where it is stated that low doses of rhBMP-2 cause formation of cartilage, which is later replaced by bone, while larger doses result in more extensive bone formation and earlier osteoinduction via intramembranous ossification. Therefore, it is possible that in the 3 d samples fractions of rhBMP-2 had been partially inactivated during the pre-incubation period, due to storage at reduced tem-peratures and in the physiological environment of the hydrogel. Presence of less active rhBMP-2 in these scaffolds resulted in less bone formation.

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Figure 2.18. Immunohistochemical staining of the 3 d samples, showing the pres-ence of cartilage-specific markers – aggrecan (left) and collagen type II (right). The cartilage tissue (C) was located in proximity to the hydrogel (G) residue. Scale bar: 100 µm. Another explanation for the endochondral ossification is that an rhBMP-2 concentration gradient was created in the highly crosslinked network of the 3 d samples, with decreasing concentrations going from the middle of the hydrogel towards the in the periphery. Hence, the peripheral rhBMP-2, com-ing in contact with MSCs, would promote the formation of cartilage, which with time would be transformed into bone. Therefore, it is possible that the entire graft would gradually be filled with bone. This means that even the 3 d hydrogels could induce improved bone formation, provided that the MSCs can enter the hydrogel network and not be prevented by the formation of a dense bone shell around the graft. However, to verify whether this hypothe-sis is true, longer in vivo experiments are required.

Figure 2.19. A schematic comparison of tissue formation induced by hydrogels prepared by different handling methods. Cartilage (C), fibrotic tissue (F), hydrogel (G) and areas containing a mixture of bone, bone marrow, hydroxyapatite and red cells (HAP) are indicated.

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The aforementioned morphological differences of the grafts from the 1 min, 5 h and 3 d groups are summarized schematically in Figure 2.19. In general, the grafts of all three groups were surrounded by a layer of fibrous tissue, which appeared to be thicker for the 1 min samples than for the 5 h and 3 d grafts. It was also noted that less cell-free hydrogel remained inside the graft, when it was surrounded by either a thin or poorly interconnected bone shell. This probably meant that this shell protected the hydrogel from cell infiltra-tion, hence slowing down the bone formation process.

Furthermore, the most important properties of the hydrogel and of the formed tissue from each test group were scored in Table 2.2. Based on this scoring, it can be said that the 1 min pre-incubation resulted in loosely cross-linked hydrogels with poor mechanical properties. Due to the partial damage of these hydrogels by the mechanical load of the surrounding tissue, less rhBMP-2 was retained. However, since the retained protein was still very active, these hydrogels gave rise to little, yet dense bone via intramembran-ous ossification.

The 5 h pre-incubation for resulted in moderately crosslinked hydrogels, which, unlike the 1 min hydrogels, could keep their integrity, despite the mechanical load of the surrounding tissue, and thus they occupied a larger volume that the 1 min hydrogels. This probably allowed for more rhBMP-2 to be retained at the site of injection. Because of this, the 5 h hydrogels in-duced bone formation with both a larger BMV and a higher BMD via intra-membranous ossification.

Table 2.2. The most important hydrogel and tissue properties, based on the experi-mental findings in Paper IV (+++ highest score, lowest score).

Properties of the hydrogel 1 min 5 h 3 d

rhBMP-2 retention ++ +++ +++ rhBMP-2 activity +++ +++ ++ Mechanical stability of the hydrogel + ++ +++ BMD +++ +++ ++ BMC ++ +++ +++ BMV + ++ +++ Fibrotic capsule ++ + + Intramembranous ossification + + Endochondral ossification +

Finally, pre-incubation for the 3 d samples resulted in the most tightly cross-linked hydrogels. The fragments, created during the extrusion of the hydro-gel through the needle during the injection, appeared to fuse together to form a compact scaffold. The mechanical stability of the hydrogel gave rise to a larger BMV than in the 1 min and 5 h grafts. However, the possible decrease in bioactivity of rhBMP-2, induced by the 3-day incubation at physiological

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pH inside the hydrogels, led to bone formation via the endochondral ossifi-cation pathway.

Naturally, the hydrogel degradation and tissue formation may be different in humans than in rats. However, based on the results of Paper IV it can be stated that different pre-incubation times give rise to materials with different mechanical and biological properties. Although all test groups resulted in formation of cancellous bone in rats, the observed variation in mechanical properties may allow physicians to adjust the treatment for each specific indication, e.g. 1 min hydrogels may be more suitable for the treatment of closed fractures, where tiny bone fragments require to be fused together, while the 5 h and 3 d hydrogels may be more preferable during treatment of large defects, where formation of maximum bone volume is of interest. However, as already mentioned, the current scenario only applies to the situ-ation observed in rats and further investigation is required in order to deter-mine what would happen should the material be used in humans.

The role of the hydrogel surface area on bone formation was further in-vestigated in Paper V, where ectopic bone formation induced by subcuta-neously implanted solid and crushed hydrogels was compared after four weeks in vivo, as described in Section 2.2.1. A morphological difference between the explanted grafts of each test group could be determined by pal-pating the specimens. The grafts of the solid hydrogel group had a hard shell and a soft core, while the grafts of the crushed hydrogel group appeared as compact discs with a homogenous stiffness. In radiographs the solid hydro-gel grafts appeared as homogenous smooth discs, while the crushed hydro-gels had a more rough and unorganized morphology (Figure 2.20).

Figure 2.20. Radiographs of explanted grafts four weeks following subcutaneous implantation of solid hydrogels (left) and crushed hydrogels (right) in Paper V, showing morphological differences between the treatments. The grafts of the solid hydrogel group appeared as smooth discs, while grafts of the crushed hydrogel group had a more unorganized and rough morphology.

Further analysis of the grafts by pQCT showed a significantly higher BMD density formed in the crushed hydrogels, compared to the solid hydrogels. On the other hand, solid hydrogels showed a trend of higher BMV than the

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crushed hydrogels (Figure 2.21). Closer examination of the quality of bone revealed that solid hydrogels induced the formation of more trabecular bone, while crushed hydrogels gave rise to more compact bone. Furthermore, the pQCT settings were such, that HAP residues were interpreted as trabecular bone.

Figure 2.21. Results of pQCT analysis of bone formation in response to hydrogels, loaded with rhBMP-2 and HAP, studied in Paper V. a) BMD, b) BMC and c) BMV presented as means ± standard deviation for n = 11 (solid hydrogel) and n = 10 (crushed hydrogel). Statistical significance: * p < 0.05; ** p < 0.01. The quality of bone was further evaluated by distinguishing between trabecular bone (280 mg/cm3 < BMD < 400 mg/cm3) and cortical bone (BMD > 400 mg/cm3). The crushed hydrogels showed significantly higher total bone mineral density and a trend of higher total bone mineral content, while solid hydrogels showed a trend of a high-er total bone mineral volume. Solid hydrogels contained more trabecular bone with a significantly larger volume of trabecular bone, and a trend of higher mineral content of the trabecular bone. The control hydrogel was detectable in the pQCT with a density similar to trabecular bone.

Histological analysis confirmed the soft bone-free core of the solid hydrogel grafts (Figure 2.22). Large spherical residues of intact hydrogel, found in the core of these samples, were surrounded by a dense shell of mineralized bone. This shell, just as it was observed in Paper IV, prevented the degrada-tion of the hydrogel, as well as cell infiltration. In general, only a few cells were observed in the grafts of the solid hydrogel group. Furthermore, the presence of dense bone and of bone formation was seen throughout the com-pact structure of the grafts of the crushed hydrogel group. The periphery of the trabecular bone was lined with osteoblasts. Osteocytes were found inside the newly formed bone matrix. The bone marrow of the crushed hydrogel grafts was filled with hematopoietic cells and adipocytes. Grafts from both test groups were surrounded by a thin layer of connective tissue.

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Figure 2.22. Histological evaluation of solid hydrogel grafts and crushed hydrogel grafts after four weeks in vivo, stained with a), d) hematoxylin/eosin and b), c), e), f) Masson’s Trichrome. a–c) In solid hydrogels, a dense bone shell surrounded a layer of adipocytes of the bone marrow and a large hydrogel residue. d–f) Crushed hydro-gels gave rise to bone formation throughout the whole sample, comprised of a mix-ture of dense bone, trabecular bone and hematopoietic cells of the bone marrow.

The data showed that the large surface area and the weak interactions be-tween the fragments of the crushed hydrogels facilitated cell infiltration, driven into the graft by the rhBMP-2 gradient. This resulted in the formation of more bone throughout these grafts, in comparison to the solid hydrogels that stimulated bone formation only on the surface of the implant.

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3 Concluding remarks

Implant loosening and severe bone damage are two major challenges in or-thopedics and dentistry. In this thesis several strategies, employing basic principles of biology, chemistry and materials science, are described as poss-ible treatments for the aforementioned conditions. The aim of this work was to develop biomaterials that would stimulate the healing of bone tissue in a controlled and reproducible way. First, the cellular response on crystalline TiO2, HAP, and HAP loaded with rhBMP-2, were characterized in vitro. The intention was to use the surfaces to create functional multilayered implant coatings that would stimulate both immediate and long-term osseointegra-tion. The coatings showed great potential in regard to promoting cell attach-ment, growth and differentiation. Furthermore, it was shown that rhBMP-2 could be loaded in the coatings by a simple soaking procedure, due to a poss-ible HAP–BMP-2 interaction.

For treatment of severe bone damage a strategy involving injectable HAA/PVAH hydrogels was employed in combination with rhBMP-2 and HAP. In order to develop a reliable and reproducible system, methods for hydrogel preparation, as well as the protein handling and detection, were first optimized. The results showed that mixing was a key parameter for the stability of the formed hydrogels. While insufficient mixing resulted in weak and not fully crosslinked hydrogels, excessive mixing led to the breaking of formed crosslinks and other chemical bonds. It was concluded that to obtain well-crosslinked hydrogels it is necessary to mix the HAA and PVAH com-ponents back-and-forth approximately 30 times. Furthermore, to avoid ex-cessive swelling and hydrogel dissolution during in vitro release experi-ments, it is advisable to use PBS as the release medium, and to allow the hydrogels to set at 37 °C for at least 3 h prior to its addition. Moreover, to prevent hydrogel dissolution, it was found necessary to always first bring the release medium to the same temperature as the hydrogel.

The handling of rhBMP-2 was also shown to be important, since the pro-tein could easily aggregate and adsorb to surfaces of tubes, pipettes and sy-ringes. A number of different types of tubes were tested, and LoBind tubes were selected for the in vitro release experiment, since they showed the least affinity for rhBMP-2. A detection method based on radioactive labeling of rhBMP-2 with 125I was used, as it allowed measuring both the released and the retained protein in vitro, which was not possible by using biochemical technique such ELISA. By using 125I labeling, rhBMP-2 retention could also

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be followed in vivo. Therefore, once the methods regarding hydrogel prepa-ration and rhBMP-2 handling and detection were established, the protein release was subsequently studied in vitro and in vivo and correlated to the protein activity and bone formation in an intramuscular ectopic model in rats. The results showed that a biphasic release profile of the active protein, where ~16 % and ~3 % of rhBMP-2 remained inside the hydrogel after 28 d in vitro and in vivo, respectively. This was significantly more than what was earlier detected by ELISA. The initial fast release phase corresponded to the early ectopic bone formation after 8 d in vivo, suggesting that the biphasic release was favorable for the recruitment of MSCs to the designated site and their differentiation into osteoblasts.

The hydrogel was additionally improved by incorporation of HAP powd-er. Once again, the interaction between HAP and rhBMP-2 was presumed to be a factor. Furthermore, bone formation could be increased by pre-incubation of the premixed hydrogel components inside the syringe prior to subcutaneous injection in rats. The difference in mechanical properties be-tween the slightly, moderately and fully crosslinked hydrogels could be use-ful, since it could permit the tuning of the treatment for each specific indica-tion. While the more compliant mechanical properties of slightly crosslinked hydrogels may be more suitable for the treatment of closed fractures, where fusion of many bone fragments is desired, the moderately and fully cross-linked hydrogels may perform better during treatment of large bone defects, where formation of maximum bone volume is of interest.

Finally it was shown that by increasing the surface area of the hydrogels it was possible to further improve their performance. Crushed hydrogels in-duced notably more bone formation than solid hydrogels, when the two types were implanted subcutaneously in rats. Moreover, there was a clear morphological difference between the grafts of the two groups. While crushed hydrogels gave rise to uniform bone formation throughout the entire graft, solid hydrogels induced formation of a dense bone shell around an intact hydrogel residue. Therefore it is concluded that increased surface area of the crushed hydrogels facilitated cell infiltration, thus improving bone formation.

The obtained information and the developed methods were subsequently used to investigate the effects of mechanical properties and surface area on growth factor release and bone formation, in the hope to more effectively use the administered dose of rhBMP-2, as well as to employ the hydrogel prop-erties to fit specific medical indications. Both the multilayered functional surfaces and the osteogenic injectable hydrogels have shown great potential for improving osseointegration and treating severe bone defects, respective-ly. Although many important parameters have been identified and optimized, many questions remain to be answered before these systems are ready to be used clinically. These questions motivated further experiments, a few of which are discussed in the next section.

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4 Future prospective

First of all, it was of interest to find out whether the multilayered coatings, which showed positive cellular response in vitro, would also function in vivo. In a recently published follow-up study (Paper IX), cylindrical tita-nium samples coated with HAP were functionalized with bisphosphonates and rhBMP-2. The functionalization was performed in the operating room, just prior to surgery, by a simple soaking procedure, similar to that described in Paper I. The samples were subsequently implanted in the distal femoral epiphysis of sheep and retrieved after six weeks. The histological analysis showed good osseointegration for all implants. Implants functionalized with rhBMP-2 induced the highest bone/implant contact compared to the other implants. However, large variations were observed for this group. This could be due to the fact that during the loading of the growth factor, regular Ep-pendorf tubes were used. Furthermore, the time of the loading varied de-pending on how long it took to anesthetize the animals. Despite these varia-tions, the study shows promising results for these coatings, giving the physi-cians the option of patient-specific functionalization of implants.

In order to better understand the HAA/PVAH hydrogels, the nature of the interaction between HAA and rhBMP-2 is currently under investigation. For this a mixture of the protein and the polysaccharide is analyzed by gel mobil-ity shift assay (GMSA); an electrophoresis technique commonly used for the study of interaction between proteins and large molecules, such as nucleic acids, i.e. DNA and RNA. The use of this technique to describe the interac-tion between synthetic dextran derivatives and rhBMP-2 has recently been described in literature. It was shown that when the rhBMP-2 interacted with the synthetic dextran, the growth factor migrated with the negatively charged polymer towards the anode, while limited migration occurred for the un-bound protein.106 The results of preliminary GMSA experiments, performed on an HAA/rhBMP-2 mixture, and compared to the native HA/rhBMP-2 mixture and rhBMP-2 alone, revealed a similar trend (Figure 4.1). HAA appeared to pull the protein more than native HA. However, large variations were observed with modification of the assay running time, current, tem-perature, ionic strength and pH. Therefore exact protocols need to be devel-oped, prior to drawing any conclusions.

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Figure 4.1. Preliminary results of gel mobility shift assays, performed on a 0.8 % agarose gel at pH 6.8, 300 mA and +4 °C for 2.5 h and 4 h, respectively, show an interaction between the HAA component of the hydrogel and rhBMP-2. The result-ing Western blots show the migration of rhBMP-2 alone as well as when mixed with native HA or with HAA. It appears that HAA binds fractions of the growth factor and pulls it towards the anode.

One of the strategies for optimization of the performance of the hydrogels involved improvement of bone formation by stimulating vascularization, via the delivery of basic fibroblast growth factor (bFGF) in addition to rhBMP-2. The idea was that by stimulating blood vessel formation, we could im-prove the cell survival inside the grafts, preventing the formation of bone only as a peripheral shell. For this rhBMP-2 was loaded into gelatin micro-spheres, which were then added to one of the hydrogel components, which also contained bFGF. The aim was that the resulting construct (Figure 4.2a) would first release the bFGF, stimulating blood vessel formation, and the-reafter the rhBMP-2, causing formation of bone. To our surprise, the intra-muscular injection of these hydrogels in rats resulted in bone formation only in the control group (Figure 4.2b) containing only rhBMP-2, and no bone was formed in the dual growth factor group (Figure 4.2c).

To investigate what caused this phenomenon it was first determined whether the bFGF was retained prior to the rhBMP-2. Surprising, it turned out that bFGF was retained better by the hydrogel system, than rhBMP-2 (Figure 4.2d), despite the fact that the latter growth factor was loaded into gelatin microspheres. This additionally supported the theory that the hydro-gel interacts with proteins through electrostatic interaction. Due to the even more basic pI of ~9 of bFGF, the hydrogels should retain this growth factor better. This data meant that the hydrogels is a promising delivery system for bFGF, for tissue engineering applications where vascularization is required.

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It also appeared that other materials were required to achieve sequential re-lease of bFGF and rhBMP-2 in the desired order.

The question remained, however, why no bone formed at all in the dual growth factor group. To investigate this phenomenon, an ALP assay was performed with constant amounts of rhBMP-2 and increasing amounts of bFGF. The results showed that bFGF inhibited the expression of ALP in vitro at a concentration a hundred times lower than rhBMP-2 as shown in Figure 4.2d. The results showed that bFGF at low concentrations may inhi-bit the differentiation of mesenchymal stem cells that is promoted by BMP-2. Therefore, a different angiogenic growth factor, which would does not inhibit the action of rhBMP-2, could be more effective.

Figure 4.2. a) Hydrogel containing bFGF and gelatin microspheres (G), loaded with rhBMP-2. Radiographs 4 weeks after intramuscular implantation of b) control hy-drogels with only rhBMP-2 and c) hydrogels with both bFGF and rhBMP-2; d) the release of rhBMP-2 and bFGF from to hydrogel/gelatin hybrid system; e) ALP ac-tivity of W20-17 cells, cultured with 100 ng/ml rhBMP-2 and 0–100 ng/ml bFGF.

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5 Acknowledgements

And now for the part of the thesis that I have been most eager to write, and which coincidently happens to be the part that normally attracts the most readers…

It is funny, how things never turn out to be the way you planed. In school, my least favorite subject was Chemistry. Logical subjects, such as Maths and Physics, always made more sense, but my dream was to become a Medical Doctor. By chance, I ended up becoming an engineer in Molecular Biotech-nology, thanks to my mom, who highlighted the program in my university catalogue and begged me to apply in case I didn’t get into medical school. For a long time, I was convinced that a PhD was not for me. However, after being introduced to the world of drug delivery and tissue engineering during my master’s thesis, I decided that I could perhaps stretch it as far as doing a PhD, with the condition that it would not be about some boring protein. Here I am five years later, an expert on one protein (BMP-2) and about to become a doctor, not in medicine, but in Polymer Chemistry. How did this happen?

The road was long and challenging, but it was far from boring! I was in-credibly lucky to end up with a stimulating and exciting project and to re-ceive support from many wonderful people, who patiently followed me on this journey. I would like to take the opportunity express my gratitude to them. In case I forget to mention someone – Thank you!

First of all I would like to thank my scientific advisors, colleagues and work friends: Jöns, for giving me the opportunity and the freedom to devel-op my PhD project, as well as taking me along on international conferences, introducing me to leading scientists and instructing me on the best skiing wax. Tim for all the Monday-therapy sessions, for your wonderful humor, support and guidance that helped me set up reasonable goals. Go brain-maps! Also I am also very happy that my gentle nagging has helped you finally “tie the knot”, before Rocky! Lars, for not only helping me with ra-dioactive labeling of BMP-2, but also taking the time to thoroughly immerse yourself into my project, despite the fact that you have your own students to take care of. Thank you! Thomas, Sune, and Torbjörn for sharing your clini-cal and industrial knowledge.

My roommate Jonas, who I currently deported to be “big in Japan”, so I could get my own office during the writing of this thesis. I know you will appreciate it, so you get your very own paragraph. This thesis would not be possible without your help! Thank you for all the crossword-, Simpsons-,

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fruit-, disco-, go-to-pick-up-parcels-breaks, as well as for all the times you proof-read my work, helped me fix things in the lab or listened to me com-plaining about my failed experiments. You have been truly the best side-kick one could wish for! Soon it is your turn and I am convinced that you will exit it with stile: S.Y.G.I.G.H! I promise to share my PowerPoint-skills when it is time for your big show. For now, may there be more sherry and less “ful öl”!

Basse, for your unique humor and for teaching me the importance of tak-ing charge over my own projects. Kristoffer, for “the hydrogel”, which has been the center of my life for the past five years, for being so well-organized, for teaching me the importance of color-coding and for patiently answering my emergency 3838 phone calls! Cissi, for all your support and patience in the lab and all the laughs during conference trips. I will never forget how you armed yourself with pens to protect yourself from robbers on in Porto! Ramiro, for teaching me all I know about Rheology and Polymer Physics, for helping me organize my experiments, proof-reading manu-scripts, which you were not even a co-author for, and taking the time to drop an ironic and somewhat nerdy joke to lift my spirits. This thesis, as many other things in our lab, would never been completed without your involve-ment. My dear colleague and side-kick #2, Gry Hullsart-Billstöm, or should I say - Senorita Nulsart-Sillström, for your friendship and assistance during animal and release experiments. I am very proud to see your transformation from a confused master student into a busy PhD student with exciting projects, who simultaneously supervises three undergrads, is a wonderful mother and receives awards for best oral presentation at international confe-rences. Bravo! Two technicians made out of pure gold: Janne, for building and repairing so many things for me along the years, from rheology equip-ments to coffee tables and broken jewelry. Brittmarie, for letting me to use various equipment, for always being so nice and for being there to answer my stressful e-mails and phone calls. Far-Side comics are the best! Oommen, for the many interesting scientific discussions. Dmitri, for involving me in your publications. The next generation Ida, Sujit, Youka, Uzma and OP. I wish you best of luck in your research. I hope your time in the group will be as much fun as it was for me! Tatti, for helping me with all the paperwork and for being my extra mom for these past 5 years! You are always so hap-py, energetic and caring! We are missing you, since you changed offices, but we only act bitter because we are jealous of the other Materials Chemistry people that get all the fun with you. They better be nice! Everyone at the Department of Materials Chemistry for all the fun during Thursday coffee, institute parties and team work during many evaluations, which we were “randomly” selected to take part in. Gunnar, Unnar, Gabriella, Petra, Van-gelis, Andrea, Adrian, Björgvin, Bengt, Lennart, Ola, Ian and everyone else at Materials Physics for letting me escape “round-bottom flask” discussions during our shared lunches. Maja for your friendship and help with “milk

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duty”. Maria, Håkan, Ulrika and Johan, for involving me in the hydroxya-patite projects and helping me publish my first paper. Alejandro, Jonas, Marjam, Cecilia, Johanna, Maryam and everyone at the Materials in Medi-cine group for all the good times both at work and during go-cart and curling get-togethers. Andreas and Natalia thank you for taking such good care of the cell lab. The Fotomol group, especially Karin, Paulo, Fernando, Gunilla and Daniel, for allowing me to trusting me with expensive equipment in your lab and for taking the time to discuss my projects. I am so glad that you were just across the corridor and willing to drop everything, even during your lunch. The folks from the EuroStec project for the stimulating collabo-rations and amazing cultural experience during our shared meetings. Elena, for all the hours you invested in helping me with animal experiments, histol-ogy and data interpretation. Thank you for your friendship! You are one of a kind! I miss working with you! Ludmila for your friendship, for doing so much important and exciting work on rhBMP-2, during your PhD, for the numerous Skype calls and all your scientific and personal advice. You are the most committed, hard-working person I know and I wish you luck in your future work! I am convinced you will go far! Stefan, Jan-Åke, Farhad, Björn at MSL for helping us with the setting up of the cell lab. Ni är bäst! All the people in the building, who have helped make Ångström a dream work place for me: Kersti, Perra, Bengt, Anti, Thomas, Sören, Desirée, Ceci-lia. Irene and all the amazing cleaning staff, for your dedication and enthu-siasm, meeting me every morning with a smile, even when I had had just spilt coffee all over the newly cleaned floor.

My life would be lonely in the sole company of cells; therefore I would also like to thank my friends: Natalia G, who has been my friend for twenty long years and who has always inspired and supported my crazy choices in life. Barnen for not forgetting about me, and being there in both hard times and fun times! Erik, Anna E, Moa, Anna H och Kalle for all the Singstar moments! Nollkå for rallys, festivals and skiing trips and inspiring me to do En Svensk Klassiker. All my new cycling friends, for giving me something else to think about besides my thesis this past year. Oleg, Olga, Vladimir and Natalia P for all the Russian dinners, Heby operas, statistics programs and general encouragement.

For the past five years whenever my mom asked me when I would finally come home for dinner, the answer was always the same: “when I have de-fended my thesis”. I am very grateful to my family for waiting: Mom, to whom this thesis is dedicated, for always being extremely proud of me and giving me your never-ending love and encouragement every day. I hope to share many dinners, long walks and shopping trips with you soon! Dad, for having your office one floor up from me, for joining me for coffee, bringing me lunch boxes, printing my posters, converting my files, solving numerous complicated equations and helping me plot and analyze data, and keeping my spirits up. Eugenia for spending endless hours proof-reading my articles

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and thesis without understanding a word. ! My grandfather, who has taught me the importance of patience and persistence in scientific work, and who, despite the fact that it was not within his field, helped me design several figures, for my articles. My grandmother Alla, for teaching me to be tidy, which is so important for lab work. My grandmother Ira, for always patiently listening to me, when I am lonely in the lab during early morning measurements, and you are the only one I can call, since everyone else is asleep. Thank you Nastya, Alexey, Lena, Sasha and Ira for all the support. Anya, I am so happy that you are interested in Biology! Make sure to study hard, so that in a few years I may read your PhD thesis! Les Brouhns for not forgetting your “petite cuisine“. Thanks to you I can almost explain my work in French. Luckily, this acknowledgement is in English, so I don’t have to put that to a test. Daniel’s family for welcoming me with open arms, Mea, Uffe, Nalle, Eva, Stefan, Adriana, Tina, and all the aunts, uncles and cousins. Special thanks, for all the moral support and high spirit, to Ann, who has been my Svensk Klassiker training buddy. We did it!

Last, but not least, I would like to thank my darling fiancé Daniel, for all your love and support. Du är krysset på min karta och jag är så tacksam att jag har funnit dig! Jag älskar dig! /

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6 Svensk sammanfattning

Benvävnad är, till skillnad från andra typer av vävnad i kroppen, mycket bra på att reparera sig själv. Vid benbrott räcker det oftast med att fixera benet i rätt läge med hjälp av ett gipsförband eller något annat yttre stödförband. Men vid komplicerade frakturer eller i situationer då benet förstörts på grund av trauma eller bensjukdom krävs det yttre stimuli för att sätta fart på ben-bildningen. Det vanligaste idag är att benvävnad transplanteras från en annan plats i kroppen. Problemet är dock att det inte finns hur mycket eget ben som helst. Dessutom kan patienter som genomgår denna behandling drabbas av infektioner och smärta i operationssåret. Det är även möjligt att transplantera ben från andra individer, men det kräver att patienten äter immunhämmande mediciner resten av livet. För att kringgå dessa problem har ett nytt tvärve-tenskapligt område skapats, där läkare och forskare arbetar tillsammans för att hitta mer hållbara behandlingsmetoder för dessa patienter. Området, som i fackkretsar kallas ”vävnadsregenerering” eller ”regenerativ medicin”, kombinerar kemi, cellbiologi, materialvetenskap och medicin för att ta fram material som kan stimulera kroppen till att reparera sig själv.

I detta doktorsarbete arbete har de ovanstående principerna används för att hitta lösningar på två benrelaterade sjukdomstillstånd. Det första tillstån-det handlar om situationer där titanskruvar (s.k. tandimplantat) som används för att fixera tandproteser, lossnar under mekanisk belastning, t ex då patien-ten tuggar. Dessa implantat är oftast tillverkade av titan, som är en mycket robust och hållfast metall. I likhet med hur en vanlig betongskruv behöver en gänga för att kunna sitta fast i väggen under vikten av en tung hylla, krävs det att tandimplantatet har maximal kontakt med benvävnad för att implanta-tet ska sitta kvar även vid mekanisk belastning. Dessvärre binder inte titan ben särskilt bra och därför riskerar sådana implantat att lossna med tiden. Därför var målet med detta arbete att förbättra titanytan hos implantatet, så att den blev bättre på att binda ben både direkt och mer långsiktigt. För detta valde vi att tillverka en flerskiktad yta, där det yttre skiktet bestod av hyd-roxyapatit (HAP), ett material som påminner om benvävnadens struktur. Syftet var att benbildande celler (osteoblaster) skulle ha lättare att ”känna igen sig” och skulle således snabbare komma igång med benproduktionen. För att ytterligare påskynda benbildningen, laddades HAP-skiktet med en molekyl kallad BMP-2, genom att doppa implantatet i BMP-2 lösning. BMP-2 är en tillväxtfaktor, ett slags molekyl som berättar för celler var de ska börja producera ny benvävnad. Då HAP och BMP-2 med tiden skulle

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brytas ner av kroppen var det viktigt att se till att det underliggande skiktet hade bra benbindande egenskaper. Därför belades titanytan med ett inre skikt av ett kristallin titandioxid (TiO2). Det kristallina TiO2 ansågs vara mer gynnsam för cellbindning än ren titan, och skulle därför stimulera långvarig förbindelse mellan ben och implantat. Genom att studera hur celler växer på de olika skikten in vitro (utanför kroppen), var det möjligt att förutspå vad som skulle häda in vivo (i kroppen). Resultaten indikerade att den flerskikta-de ytan hade stor potential. Dessutom visade det sig att HAP kan binda BMP-2 och på så sätt förlänga tillväxtfaktorns verkan genom att hålla den på plats. Då BMP-2 snabbt bryts ner i kroppen var upptäckten extra viktig, då den påvisade att det gick att effektivare använda molekylen.

Det andra sjukdomstillståndet var avsaknad av ben på grund av skada el-ler sjukdom. För att kunna stimulera benbildning på en specifik plats utan att behöva transplantera ben eller genomföra komplicerade kirurgiska ingrepp, valde vi att injicera BMP-2. Återigen var BMP-2-stabiliteten ett problem och för att förlänga tillväxtfaktorns verkan injicerades den tillsammans med en syntetisk bärare framställd av två polymerer. En polymer är en lång molekyl som består av många repeterande enheter (monomerer). På samma sätt som tåg kan bestå av likadana eller olika vagnar så kan ju även polymerer göra det. Fördelen med att använda polymerer inom medicin är att dessa kan till-verkas så att de bryts ned genom naturliga mekanismer i kroppen. Genom att variera tillverkningsprocessen kan materialets styvhet och nedbrytningstid anpassas för det medicinska ändamålet. I detta arbete använde vi oss av två polymerer – hyaluron (HA) och polyvinylalkohol (PVA). HA är en polymer som består av tusentals ihopkopplade sockermolekyler och finns bland annat i bindväv hos alla människor och djur. Den har tidigare utvunnits från tupp-kammar, men kan numera framställas av bakterier. PVA är en syntetisk po-lymer som ofta används inom medicin. I ett tidigare arbete modifierades HA och PVA kemiskt så att de två polymererna kunde kopplas ihop med var-andra. Modifieringen gjordes genom att kemiskt förändra ungefär 5 % av monomererna i varje polymerkedja, d.v.s. 5 av 100 ”vagnar” i varje ”tåg” fick en krok på sidan med hjälp av vilken de kunde haka in i en annan vagn med en krok på ett parallellt tåg. Inom polymerkemi brukar man säga att polymerer på detta sätt bildar ”tvärbindningar”. När vattenlösningar av mo-difierad HA och PVA blandas börjar polymererna genast haka in i varandra och bildar ett tredimensionellt nätverk. Nätverket är en geléliknande sub-stans (s.k. hydrogel) som i stort sett består utav vatten och endast 1% av polymerblandningen.

Syftet med detta arbete var att optimera hydrogelsystemet så att det blev både enkelt att använda och fungerade som en effektiv bärare för BMP-2. Först och främst undersöktes hur man rent praktiskt skulle blanda de två polymerlösningarna för att få optimalt många tvärbindningar och på så sätt tillverka stabila hydrogeler. Därefter valdes en passande vätska för att simu-lera in vitro hur BMP-2 skulle frisättas från hydrogelerna i kroppen, utan att

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dessa skulle riskera att svälla eller lösa upp sig under experimentets gång. För att kunna följa tillväxtfaktorns frisättning, märktes BMP-2 in med en radioaktiv isotop av jod (125I). 125I-märkingsmetoden har funnits länge, men den har aldrig tidigare använts för att studera BMP-2 frisättning från HA–PVA-hydrogeler. Fördelen med att märka in BMP-2 var att man på så sätt lätt kunde mäta mängd radioaktivitet, och alltså mängd BMP-2. Med hjälp av 125I-märkingsmetoden gick det att mäta BMP-2 i hydrogelen, i vävnad, på sprutor och i provrör. Dessutom, om flera BMP-2 molekyler råkar klumpa ihop sig brukar den alternativa mätmetoden ELISA ha svårt att upptäcka tillväxtfaktorn, medan 125I-märkingsmetoden inte hade några sådana be-gränsningar. Det var alltså möjligt att följa BMP-2 både in vitro och in vivo. Frisättningsexperiment visade att BMP-2 frigörs från hydrogelen i två faser: snabbare under första veckan och långsammare därefter. Tillväxtfaktorn behöll sin biologiska aktivitet i hydrogelerna, och då materialet injicerades i lårmuskeln hos råttor bildades nytt ben redan efter åtta dagar.

En rad experiment genomfördes där varianter av hydrogelen studerades. Resultaten visade att genom att tillsätta HAP-pulver till hydrogelen kunde BMP-2 hållas kvar längre och kunde således användas på ett effektivare sätt. Dessutom var det möjligt att få mer ben att bildas med samma hydrogel och BMP-2-dos genom att blanda polymererna och därefter vänta några timmar innan injektion. Då hann fler polymermolekyler hitta varandra och bilda tvärbindningar, vilket resulterade i stabilare hydrogeler som lättare kunde hålla sin form i kroppen. Slutligen kunde det påvisas att det, genom att kros-sa den förblandade hydrogelen en dag efter blandningen, gick att skapa ma-terial med större ytarea, vilket gjorde det lättare för celler att komma in i materialet och bilda benvävnad.

I detta doktorsarbete beskrivs således metoder för att stimulera bildning av benvävnad på titanimplantat samt vid svåra benskador. Både de flerskik-tade titanytorna och hydrogelerna är lätta att framställa och använda. De gör det möjligt för kirurgen att anpassa behandlingen till det specifika patientfal-let. Det finns dock många faktorer som återstår att studera innan dessa meto-der kan användas på patienter, så att behandlingen blir säker, kostnadseffek-tiv och reproducerbar.

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