Designing and Constructing a Prototype Standalone Bioreactor

Designing and constructing a prototype standalone bioreactor using 3D printing and finite element analysis: A tool to define osteogenic differentiation in a 3D mechanical environment

Sigurður Rúnar Rúnarsson

Thesis of 60 ECTS credits submitted to the School of Science and Engineering at Reykjavík University in partial fulfillment of the requirements for the degree of Master of Science (M.Sc.) in Biomedical Engineering

January 2019

Supervisors: Ólafur Eysteinn Sigurjónsson, PhD Professor, Reykjavík University, Iceland

Paolo Gargiulo, PhD Associate Professor, Reykjavik University, Iceland

Examiner: Hans Guttormur Þormar, Examiner CEO

Designing and constructing a prototype standalone bioreactor using 3D printing and finite element analysis: A tool to define osteogenic differentiation in a 3D mechanical environment

Sigurður Rúnar Rúnarsson

January 2019

Abstract Tissue engineering aims to develop methods to construct tissues in vitro that have identical tissue-specific morphological, biological, chemical, and mechanical properties to those cultured in vivo. In vivo tissues are embedded into complex environments, communicating both with nearby zones and with the whole organism, which determines the tissue-specific function. Special instruments called bioreactors are used to mimic these conditions outside of the body. Bioreactors provide a controlled environment where specific parameters can be determined by the researcher, to match desired biological condition. In this project, an existing bioreactor system was redesigned and improved using computer aided design and 3D printing technology to make it operational for future studies involving osteogenic differentiation. This bioreactor system provides compression on scaffolds located in a chamber. It also has indirect perfusion, maintains a temperature of 37°C when running, and can maintain an optimal pH value of (7.2 to 7.6) for osteogenic differentiation. The main issues in the previously-existing system were: sterility problems; bulging of the bioreactor chamber during heating, leakage of culture media along joints during bioreactor operation; and lack of pH stability. Several iterations of changes by trial and error were made to improve the overall design. Using µCT technology, a screw mechanism was attached to the chamber which proved to be an important addition. An internal plate system was implemented to ease scaffold placement among other uses. Further ideas for future development were also discussed. A manual was made for the bioreactor and protocols were developed for cleaning and sterilizing the chamber. Bioactive scaffolds seeded with mesenchymal stem cells were used in test experiments. Additionally, a finite elements analysis was carried out on the bioreactor chamber to quantify the compression and perfusion speed necessary for viability of mesenchymal stem cells inside the bioreactor system. A flow analysis of culture media was also analyzed for different chamber designs.

Þróun og smíði á frumgerð lífferlakerfis með þrívíddarprentun og einingaaðferðargreiningu: Tæki fyrir beinsérhæfingu í þrívíddar mekanísku umhverfi

Sigurður Rúnar Rúnarsson

Janúar 2019

Útdráttur Vefjaverkfræði miðar að þróun á aðferðum við að smíða vefi utan líkama sem hafa sömu eiginleika í vefjaformgerð, lífræðilega, efnislega og mekaníska og þeir vefir sem vaxa innan líkamans. Innvortis vefir eru umvafðir flóknu umhverfi og eiga samskipti við nágrannasvæði sem og við alla lífveruna sem ákvarðar eiginleika vefjarins. Lífferlakerfi eru notuð til að líkja eftir innvortis aðstæðum utan líkamans. Lífferlakerfi bjóða upp á stýranlegt umhverfi þar sem færibreytur geta verið aðlagaðar til að líkja sem best eftir ákveðnum aðstæðum. Í þessu verkefni var tekið fyrir lífferlakerfi í mótun sem þarfnaðist endurhönnunar og bætingar. Tölvuteikningarhönnun og þrívíddarprentunartækni var beitt til þess að koma lífferlakerfinu í rekstur fyrir komandi rannsóknir sem snúa að beinsérhæfingu. Þetta lífferlakerfi gefur kost á samþjöppun á burðarvirkjum frumna, staðsettum í sérhönnuðu íláti Kerfið býður einnig upp á gegnumflæði sem viðheldur 37°C við keyrslu, og heldur ákjósanlegu sýrustigi (7.2 – 7.6) fyrir beinsérhæfingu. Helstu endurbætur frá upprunalegu útgáfunni voru: Einangrunarhæfni; endurbót á íláti vegna hitunaráhrifa, endurbót á íláti vegna leka; og prófanir á sýrustigsstýringu. Nokkrar ítranir af breytingum voru gerðar til þess að bæta heildarhönnunina. Með hjálp tölvusneiðsmyndtækni bættist inngreyptur skrúfgangur við ílátið sem reyndist góð og mikilvæg viðbót. Plötukerfi í íláti var bætt við til þess að auðvelda handtök við að staðsetja burðavirki, auk fleiri hlutverka. Hugmyndir um frekari framhald á hönnun eru til umfjöllunar. Leiðarvísir var búinn til fyrir lífferlakerfið sem og leiðbeiningar fyrir dauðhreinsun á ílátinu. Einnig voru gerðar leiðbeiningar fyrir hreinsun á boxinu eftir þrívíddarprentun. Lífvirk glerburðarvirki sáð með mesenkímal stofnfrumum voru notuð í prófunarkeyrslum. Loks var einingaaðferð beitt á boxið til að magngreina samþjöppunina og gegnumflæðishraða sem hentaði mesenkímal stofnfrumunum í lífferlakerfinu. Flæðisgreiningu á frumuæti var beitt fyrir mismunandi hönnun á ílátinu.

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Designing and constructing a prototype standalone bioreactor using 3D printing and finite element analysis: A tool to define osteogenic differentiation in a 3D mechanical environment

Sigurður Rúnar Rúnarsson

January 2019

Student:

Sigurður Rúnar Rúnarsson

Supervisors:

Ólafur Eysteinn Sigurjónsson

Paolo Gargiulo

Examiner:

Hans Guttormur Þormar

The undersigned hereby grants permission to the Reykjavík University Library to reproduce single copies of this Thesis entitled Designing and constructing a prototype standalone bioreactor using 3D printing and finite element analysis: A tool to define osteogenic differentiation in a 3D mechanical environment and to lend or sell such copies for private, scholarly or scientific research purposes only.

The author reserves all other publication and other rights in association with the copyright in the Thesis, and except as herein before provided, neither the Thesis nor any substantial portion thereof may be printed or otherwise reproduced in any material form whatsoever without the author’s prior written permission.

date

Sigurður Rúnar Rúnarsson Master of Science

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I dedicate this to all the incredible people surrounding me and supporting me through times I thought I could not do it. My parents and friends who have believed in me from day one. My supervisors for keeping up good spirit and guiding me.

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Acknowledgements

Strength does not come from winning. Your struggles develop your strengths. When you go through hardships and decide not to surrender, that is strength. Arnold Schwarzenegger

I want to thank my supervisors, Ólafur Eysteinn Sigurjónsson and Paolo Gargiulo, for everything they have done for me this past year and for giving me the opportunity to continue development of the bioreactor system. It has been a huge step for me getting to know fundamental aspects of tissue engineering and to be a part of great groups at Reykjavik University and the Blood Bank. There were many people involved in making this happen. First, I want to thank Joseph Lovecchio for allowing me to continue his work, for helping me in the development of the bioreactor system, and for his great hospitality when I visited him in his university (University of Bologna) for two weeks in Cesena, Italy. Also, thanks to Gissur Örlygsson for his much-appreciated work involving the µCT scanner. I want to thank, in alphabetical order: Jordyn Chapman, Christian Christensen, Kasper Bruun Løvenstein, Marta Mikaelsdóttir, Helena Montazeri, Francesca Picco, Þóra Björg Sigmarsdóttir, and all other people involved in my project at the Blood Bank and Heilbrigðistæknisetur at Reykjavik University. Their help was unselfish, and I am very thankful for it. Finally, I want to thank my parents: María Antonsdóttir and Rúnar Óskarsson for giving me mental support during my five years of studying at Reykjavik University.

This work was supported by RANNIS grant nr. 174398-05. Reykjavik University supported my trip to Italy. Also, all the 3D material used for developing the bioreactor chamber was provided by Heilbrigðistæknisetur at Reykjavik University.

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Contents

Acknowledgements ...... xv

Contents ...... xvii

List of Figures ...... xix

List of Tables ...... xxi

List of Abbreviations ...... xxiii

1 Introduction ...... 1 1.1 Regenerative medicine ...... 1 Tissue Engineering (TE) ...... 3 Bone tissue engineering ...... 4 Scaffolds in tissue engineering ...... 5 Seeding of scaffolds ...... 6 3D printing scaffolds ...... 6 Stem cells ...... 8 Embryonic stem cells (ESCs) ...... 9 Somatic stem cells ...... 9 Mesenchymal stem cells (MSCs) ...... 10 MSCs tri-lineage differentiation ...... 10 Induced pluripotent stem cells ...... 11 2D vs 3D cultures ...... 11 Bioreactors ...... 13 Properties of bioreactors ...... 13 Bioreactor materials ...... 15 Bioreactor systems in tissue engineering ...... 15 Systems using hydrodynamic shear stress ...... 15 Perfusion-based bioreactor systems: ...... 17 Mechanical strain bioreactor systems: ...... 18 Electromagnetic field bioreactors: ...... 19 In vivo bioreactors ...... 19 3D printing of bioreactor components ...... 20 Computer drawings ...... 20 Microcomputed tomography (µCT) ...... 20 µCT in regenerative medicine ...... 21 Finite elements analysis ...... 21 Wall shear stress ...... 22 Perfusion ...... 22 Compression ...... 23 xviii

Aims and Objectives ...... 24

2 Background ...... 25

3 Methods & Results...... 26 Bioreactor evaluation and redesign...... 26 3D printing ...... 26 Screw mechanism addition ...... 26 Spiros and the Autoclave connectors ...... 28 Screw port tests ...... 28 Chamber version development ...... 30 Version 1 ...... 30 Version 2 ...... 31 Version 3 ...... 32 Version 4 ...... 32 Final chamber configuration ...... 34 Professional drawings using Inventor ...... 34 Finite element analysis...... 38 Meshing and preprocessing ...... 38 Compression ...... 40 Parameters ...... 40 Shear stress ...... 41 Flow analysis ...... 42

4 Discussion ...... 45 Bioreactor...... 45 Bioreactor status ...... 45 Future improvements ...... 46 Existing bioreactors ...... 48 Finite element analysis...... 50 Conclusion ...... 51

5 Bibliography ...... 52

Appendix A – Bioreactor manual ...... 59

Appendix B – Cleaning of chamber for bioreactor system (at RU) ...... 76

Appendix C – Cell experiments in bioreactor and other ...... 80

Appendix D – Matlab Code...... 85

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List of Figures

Figure 1.2 Differentiational ability of mesenchymal stem cells [41] ...... 10 Figure 1.3 Paradigm of TE: ...... 14 Figure 1.4 The spinner flask bioreactor [17] ...... 16 Figure 1.5 A rotating bioreactor system: A) scaffolds are attached to the outer vessel wall during rotation; B) rotating bed bioreactor; C) Cells are seeded on discs rotating on the horizontal axis [17] ...... 17 Figure 1.6 Schemetic drawing of perfusion flow principles. A) shows indirect perfusion where the medium flows around the scaffold, while B) shows direct perfusion, where medium is forced through the scaffolds. Shear stress is therefore transferred directly to the cells within the scaffolds. Arrows represent magnitued and direction of flow [17] ...... 18 Figure 3.1 Left: A Luer lock style connector. Right: A computer generated image of the connector on the left. This was made with a µCT scanner and Autodesk Inventor software27 Figure 3.2 Screenshot from Autodesk Meshmixer when scew ports (grey, lime, purple, green and cyan colored) were inserted onto the chamber ...... 27 Figure 3.3 Left: A Spiros® connector. Right: A Microclave® connector. They are both made by ICU Medical ...... 28 Figure 3.4 Screw test strand (left). Testing the screw ports for strength and leakage (right) 29 Figure 3.5 Comparison between chambers before starting redesigning (right) and after improvements were made (left) ...... 30 Figure 3.6 Version 1; connectors too close together ...... 31 Figure 3.7 Version 1 chamber (left) fully cleaned and version 2 chamber (right) straight out of the printer. The yellow tint is caused by UV radiation and can be reduced significantly by using light treatment (Appendix B) ...... 31 Figure 3.8 Addition of pipe system connecting two far right screw ports down to the culture medium ...... 32 Figure 3.9 The internal plate (left) and the supporting plate (right). Further component description can be found in Appendix A – components overview ...... 33 Figure 3.10 Empty chamber (left) and chamber with the internal plate system (right) ...... 33 Figure 3.11 Internal plate and supporting plates in a hydrogel molding configuration. The molding takes place in between teeth near the bottom ...... 34 Figure 3.12 A schematic view of the port hole purposes of the bioreactor chamber...... 34 Figure 3.13 Schematic drawing of the bioreactor chamber made in Inventor. Dimensions are in millimeters...... 35 Figure 3.14 Schematic view of the bioreactor chamber including the internal plates...... 36 Figure 3.15 Schematic view of other chamber components. Dimensions are in millimeters...... 37 Figure 3.16 Geometry representing liquid media space inside the bioreactor chamber ...... 38 Figure 3.17 Histograms for differently detailed meshes of the media geometry ...... 39 Figure 3.18 Computer generated mesh of the media area of the chamber ...... 39 Figure 3.19 Young’s modulus, Poisson’s ratio and density as a function of porosity. The y coordinates seen is the graphs, are the obtained values for the bioactive glass scaffolds used. The Matlab code can be found in Appendix D ...... 41 Figure 3.20 The shapes of chamber used for FE analysis, a) Original chamber, b) chamber xx with tapered edges, c) chamber with rounded edges ...... 42 Figure 3.21 Total velocity flow in meters per seconds in all directions at inflow level ...... 42 Figure 3.22 Total velocity flow in meters per second in all directions, at scaffold level ...... 43 Figure 3.23 Total velocity flow in meters per seconds only in x-directions, at inflow level 43 Figure 3.24 Total velocity flow in meters per seconds only in x-directions, at scaffold level ...... 44

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List of Tables

Table 1.1 Examples of materials used in RM [8] ...... 2 Table 1.2 Three-dimensional printing techniques for bone scaffold fabrication [30] ...... 7 Table 1.3 Advantages and disadvantages for different stem cell type uses in TE [8] ...... 8 Table 1.4 Categorization of stem cells [36] ...... 9 Table 3.1 Three 3D-printed strands with nine different ports for determining the best material and thickness of screw mechanism. Materials used in test the strands were: MED610, Veroclear hard plastic, and Agilus 30 rubber. Some pieces were also coated with MED610 (adds 0.3 mm layer on the surface). The best combination is shown in bold...... 29 Table 3.2 Relevant equations as a function of porosity for mechanical parameters ...... 40 Table 3.3 Results from the FE analysis. Shear stress evaluation from compression, perfusion, and both combined...... 41 Table 4.1 Selected commercially available bioreactor systems [17]...... 48

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List of Abbreviations

2D Two-dimensional 3D Three-dimensional AM Additive Manufacturing ASC Adult Stem Cell BTE Bone Tissue Engineering CAD Computer Aided Design CT Computed Tomography DICOM Digital Imaging and Communications in Medicine DNA Deoxyribonucleic Acid ECM Extracellular Matrix EMF Electromagnetic Field ESC Embryonic Stem Cell FDM Fused Deposition Modeling FEA Finite Element Analysis GMP Good Manufacturing Practice ICM Inner Cell Mass iPSC induced Pluripotent Stem Cell MSC Mesenchymal Stem cell MSc Master of Science NASA National Aeronautics and Space Administration PEMF Pulse Electromagnetic Field PhD Doctor of Philosophy PLA Polylactic Acid RM Regenerative Medicine RU Reykjavik University STL Stereolithography TE Tissue Engineering TGF-β Transforming Growth Factor beta USNCB United States National Committee UV Ultraviolet µCT Micro-Computed Tomography

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11 Introduction

1.1 Regenerative medicine

From beginning of time, humans have sought ways to improve their existence. Diseases, injuries, and congenital malformation have always been a part of the human experience. Imagine if damaged bodies could be restored biologically: life could go on as though tragedy had not intervened. Artificial and prosthetic materials for replacing limbs, teeth, and other tissues results in a partial restoration of lost function, and we have already seen extraordinary achievements in that field [1]. Athletes have even been denied participation in the Paralympics because they cannot prove that their prosthetics do not give them an unfair advantage [2]. In the near future, the goal is to add a better way to solve the problem of restoring lost function, a more biological approach; this is where regenerative medicine comes in. Regenerative medicine (RM) involves the replacement or regeneration of human cells, tissues, or organs to restore or establish normal function [3]. The term itself, regenerative medicine, is widely considered to have been originated by William Haseltine during a conference on Lake Como in 1999, in his attempt to describe an emerging field, that blended knowledge derived from different subjects such as tissue engineering, cell transplantation, stem cell biology, biomechanics prosthetics, nanotechnology, and biochemistry. “In the future, many new medicines will be human substances. Those substances are mostly interchangeable among the members of our species. … Human substances used as drugs have the ability to repair, rebuild and restore injured, diseased or worn-out tissues. It is conceivable that over time we will gain enough information to control the behavior of every cell in our bodies. Once we have achieved such mastery, we will be able to heal any disease” [4]. Dr. Haseltine also recognized cells that had the unique ability to differentiate into other types of cells. Regeneration of body parts is a common phenomenon in nature. Perhaps the best-known example is that of the salamander, who can regenerate a lost tail in few days. Another example is the developing fetus, which has the ability to heal dermal skin wounds by regenerating normal epidermis and dermis with restoration of the extracellular matrix 2

architecture [5]. However, this particular repair ability is lost later in gestation. The immune system can develop chronic rejection and destruction over time if care is not taken in tissue engineering (TE) approaches, which is a major issue; however, using patient’s own stem cells eliminates this problem [1]. RM holds promise not only as a means to compensate for donor shortages, but also as a means to improve the standard of care [6]. As previously mentioned, RM is a multidisciplinary science, that has evolved in parallel with recent biotechnological advances. Material specialists can now fabricate biocompatible scaffolds with a wide range of physical parameters, combining mechanical integrity with high porosity to promote cell infiltration [7]. The field of RM can be split into three different categories. The first type of RM is cellular therapy, which consists of injecting healthy cells in pathologic tissues. The cells are either previously differentiated or undifferentiated stem cells which can differentiate in vivo depending on circumstances. A variety of cell types can be used [8]. These cell types are differentiated endogenous primary cells, adult stem cells, embryonic stem cells, induced pluripotent stem cells, and amniotic fluid-derived stem cells.

Table 1.1 Examples of materials used in RM [8]

Origin Examples Natural materials Collagen, fibrin, chitosan, dextran, alginate, gelatin, cellulose, hyaluronic acid, silk fibroin Acellular tissue matrix Bladder acellular matrix, small intestinal submucosa, bowel acellular tissue matrix, bovine pericardium, human placental membrane Synthetic polymers Polyglycolic acid, polylactic acid, poly(lactic-co-glycolic) acid, polycaprolactone, poly(copralactone-co-ethyl ethylene posphate), polydioxane, polyethylene glycol, poly-N-(2-hydroxyethyl)metacrylamide, poly-N-(2-hydroxypropyl) methacrylamide

The second approach is biomaterials. Last few years, extracellular matrices (ECM) have proven to be vital in many different functions such as gene expression, survival, death, proliferation, migration and differentiation of cells. These ECMs or scaffolds can be either classified as natural or synthetic with different advantages and disadvantages. Table 1.1 lists up materials used in RM; categorized by natural-, acellular tissue-, and synthetic materials. Synthetic materials can be reproduced on a large scale, but their main disadvantage is lacking

3 biologic recognition. Natural scaffolds can be integrated perfectly and with right methods, it can eliminate all immune response problems (by creating acellular tissue matrices [9]). The main disadvantage of natural scaffolds is that they cannot be produced easily in large quantities according to good manufacturing practice (GMP) [8]. Material used should be biocompatible and should biodegrade at the same rate as the regeneration process. The rate of biodegradability can be somewhat controlled by synthesizing the scaffold by using a chemical crosslinking agent [10]. A very important property is the pore size (porosity), which plays an important role in the flow of nutrients and waste. It can be difficult to manufacture materials with desired pore size, but three-dimensional (3D) printers provide a good solution to that problem. This will be discussed further in later chapters. The third approach to RM is a combination of the two aforementioned approaches which is called implantation of scaffolds seeded with cells. There have been a number of examples of this method over the past several years. For example, in 2006, Atala et al. reported on autologous engineered bladder constructs that could be used in patients [11]. Another example takes advantage of the decellularized liver matrix, which is recellularized and then transplanted into a human subject [10], [12]. This idea of splitting RM into cell therapy and tissue engineering is the basis of the topic of this thesis, which focuses on tissue engineering.

Tissue Engineering (TE)

As with RM, the focus of tissue engineering (TE) is restoring, maintaining and improving the function of damaged tissues and organs. Unlike RM, however, TE focuses on developing tissue outside of the body [13]. The goal of TE is to repair or replace tissues and organs by delivering implanted cells, scaffolds, DNA, proteins, and/or protein fragments at surgery. TE is a multidisciplinary field that merges engineering and biology. Material science has also become a vital part of this field due to the integration of support structures [14]. As a field, TE has been defined for three decades. Much still needs to be learned and developed to provide a firm scientific basis for therapeutic application. TE can draw on the knowledge gained in the fields of cell and stem cell biology, biochemistry, and molecular science, while chemical engineering, and bioengineering allow the rational application of engineering principles to living systems. In addition, the fields of genetic engineering, cloning, and stem cell biology may ultimately develop hand in hand with the field of tissue engineering in the treatment of human disease, each discipline depending on developments in the others [1]. There are several things to consider if replacement tissues and organs are to have the same 4

functions as their original counterparts. For example, their biomechanical properties are critical. A subcommittee was formed by the United States National Committee on Biomechanics (USNCB) to address questions such as: What are the thresholds of force, stress and strain that normal tissue encounter? What are the mechanical properties of tissues when subjected to stresses and strains, as well as under failure conditions? Which of these properties should a tissue engineer insist upon incorporating into the design? These questions fall under the concept of Functional Tissue Engineering [14]. The tissue engineering work by Vacanti et al. [15] in 1988 has inspired many important advancements in this field, bringing it closer to achieving its potential as a life-saving and life-improving option for countless patients, for whom suitable medical treatments do not yet exist. The need for success in this field is too apparent. Currently there are 114,563 candidates on waiting lists for transplants in the US alone, while the number of transplants performed between January and August 2018 were only 24,213 (http://optn.transplant.hrsa.gov). The average waiting time for kidney transplants, for example, is 3 to 5 years (with high deviations) [16]. One of the many areas of concern are bone diseases such as bone infections, bone fractures, osteoarthritis, osteoporosis, and rheumatoid arthritis, which are all problems that require surgical interventions. These diseases have become a socioeconomic problem due to increasing life expectancy [17], this embraces a new subfield of TE, called Bone tissue engineering (BTE).

Bone tissue engineering

Bone related injuries and diseases represent a significant portion of total worldwide healthcare cost. Out of the six million fractures occurring every year in the United States, 10% are complicated with impaired healing [18]. The potential of BTE to accelerate bone regeneration and decrease the number of impaired healing cases offers an exciting new option for addressing the reconstruction of osseous defects [1]. Currently, when a clinician is confronted with a large bone defect resulting from either disease or trauma, the best option for reconstructing is to use autografts because they provide osteogenic cells, osteo-inductive factors, and the lattice (scaffold) needed for bone regeneration. The autograft can be obtained from the patient’s skull, ribs, iliac crest, distal femur, or proximal tibia. The main disadvantages of this method are the relatively high risks

5 of harvest surgery, the fact that autografts are scarce resource, and the possibility that the autograft is insufficient or impossible to use. An alternative to autograft use, an allograft from donors or deceased can be used, but with allografts follows an increased risk of disease, infection, and immunologic rejection [1]. BTE is a field that aims to introduce an alternative to this conventional method. The discipline of BTE involves the use of osteoconductive matrices (scaffolds), bone-forming cells, and osteogenic growth factors [17] as a solution to bone defects. Cells are the living component of a tissue construct, and they, by themselves, can produce matrix neotissue inside of scaffolds [19]. The hope is that, In the near future, clinicians will be able to choose between these two methods, which would lower the reliance on the autologous method. In the next section, an important part of BTE is discussed: the scaffolds (the 3D matrices).

Scaffolds in tissue engineering

Scaffolds are a vital part of TE, and a number of designs have been experimentally and clinically studied. An ideal scaffold should have following properties: 1) it should be 3D and highly porous with an interconnected pore network for nutrient and waste transport; 2) it should be biocompatible and bioabsorbable with a controlled degradation and resorption rate to match tissue growth; 3) the scaffold surface must be suitable for cell proliferation, attachment, and differentiation; and 4) it should have matching mechanical properties to the tissues at sites of implantation. Scaffolds are normally made of materials like hydroxyapatite (HA), poly(α-hydroxyesters), hydrogels, natural polymers such as collagen and chitin, and bioactive glass [20]. The discovery of 45S5 bioactive glass by Hench et al. in 1971 [21] started the field of bioactive materials. Bioactive materials are intended to interact with the biological environment and to change the extracellular matrix (ECM) for desired effect. These materials are a good choice for scaffolds because of characteristic features such as time - dependency and dynamic modification of the surface when it is in contact with biological fluid [22]. An important effect of bioactive glass is its ability to control ion dissolution products. These products affect the gene expression of osteoblasts, which leads to bioactive glass having bone regenerative abilities [22]. There are many different types of bioactive glass used, each of yield different rates of in vivo bone regeneration and bone healing. 6

Seeding of scaffolds

Two important factors in BTE is number of cells used for seeding scaffolds and how the seeding is performed. Van den Dolder et al. observed higher calcium contents in rat bone marrow cells when a higher density of cells was used for seeding [23]. A minimum of 8 × 104 and an optimum of 8 × 106 bone marrow stromal cells/cm3 were determined for successful bone formation in Dutch milk goats [24]. Other factors such as seeding incubation time and scaffold morphology also affect seeding efficiency [17].

3D printing scaffolds

There are many different methods to make porous bone scaffolds, including chemical/gas foaming [25], particle/salt leaching [26], freeze drying [27], thermally-induced phase separation [20], and foam-gel [28]. Each of these methods share an important disadvantage: pore size, shape, and interconnectivity are hard to control. However, additive manufacturing (AM) approaches can overcome this hurdle. There are many types of AM approaches, and the best example is 3D printing. 3D printing technology was developed in the early 1990s at the Massachusetts Insitiute of Tehnology (Cambridge, MA) by Sachs et al. [29]. This early technology was based on having powder covered on the printing bed, then using binders for forming objects, layer by layer. This method is useful for fabricating scaffolds with tailored porosity from a computer aided design (CAD) file. Before printing, important parameters such as powder packing density, powder flowability, layer thickness, binder drop volume, binder saturation and powder wettability need to be optimized. A large variety of ceramic, metallic, polymeric, and composite materials can be processed using 3D printing. This method allows control over fine features, including interconnected porosity, and has no contamination issues. Its main drawback is the extensive optimization needed to get good quality parts [30]. Low mechanical strength is a major challenge with porous scaffolds; this is also true for 3D printed ceramic scaffolds. Table 1.2 summarizes most important aspects of 3D printing bone scaffolds.

Table 1.2 Three-dimensional printing techniques for bone scaffold fabrication [30]

Processing details Processed materials for bone TE Advantages (+) and Disadvantages (-) Strands of paste/viscous PCL (+) Mild condition of material (in solution form) process allows drug extrusion based on the Hydroxyapatite and biomolecules predesigned structure (proteins and living Bioactive glass cells) plotting Layer by layer deposition of strands at constant rate, Mesoporous bioactive glass/alginate composite (-) Heating/post- under specific pressure processing needed Polylactic acid/polyethylene glycol for some materials Disruption of strands restricts the according to the tear of PLA/polyethylene glycol/G5 glass biomolecule speed incorporation Poly(hydroxymethylglycolide-co-ɛ- caprolactone)

Bioactive 6P53B glass

Fused deposition modeling (FDM) is a commercially available 3D printing technology used in many modern 3D printers (such as Stratasys 3D printer types). This is a layer additive manufacturing process that uses production grade thermoplastic materials to produce both prototypes and end-use parts. The process begins by slicing 3D CAD data into layers. The data is then transferred to a machine which constructs the parts layer by layer. Thin thread- like spools of thermoplastic and support material are used to create each cross-section of the part. Fresh material is slowly extruded through dual heated nozzles. These extrusion nozzles lay down both support and thermoplastic material upon the preceding layers. The nozzle continues to move in a horizonal XY plane, while the build platform moves down, thus building the part layer by layer (figure 1.1). The finished part is removed from the build platform and the support material is removed. Raw FDM parts have visible layer lines. However, there are multiple finishing options to create smooth, even surfaces by hand sanding, wet sanding, or treating the material with chemicals such as salt or other solvents [31]. This type of 3D printing is very useful for many different industries, including the healthcare industry. It is possible to 3D print with biocompatible materials such as MED610 (made by Stratasys) [32], which is important for applications such as creating biocompatible housing chambers for viable stem cell growth.

Stem cells

The use of live cell-based therapies in medicine is not a new concept. The first successful allogeneic human hematopoietic stem cell transplant took place in 1968 [33], and is now a routine clinical procedure for bone marrow regeneration. Since the 2000s, there has been a steady increase in the number of cell-based therapy clinical trials, with an increasing spread of phases and number of target indications [34]. This increased diversity of target clinical indications has been largely driven by growing evidence of the trophic effects of infused cells. [35]. By definition, a stem cell is characterized by its ability to self-renew and its ability to differentiate along multiple lineage pathways. Stem cells can be classified as embryonic (ESC), fetal and adult stem cells (ASC), based on their origin. Table 1.3 lists the advantages and disadvantages of different stem cell types in TE.

Table 1.3 Advantages and disadvantages for different stem cell type uses in TE [8]

Cell type Advantages Disadvantages Differentiated - No tissue rejection - Difficult expansion because of endogenous - Reduced inflammatory response in vitro short lifespan primary cells - Difficulty in getting healthy cells in diseased organs

Adult stem cells (ASCs) - No tissue rejection - Low number of cells - No ethical problems in each tissue - No tumors - Difficult in vitro expansion - Easy isolation without differentiation - In some cases easy access Embryonic stem cells - Unlimited ability to self-renew - Ethical and political problems (ESCs) - Potential to differentiate into many specialized - Tumorigenicity cells from all the three germ layers - Need for feeder cell layers Induced pluripotent - Similar as ESCs - Tumorigenicity stem - Easier generation than ESCs cells (iPSCs) - No ethical problems

Amniotic fluid-derived - No tumorigenic - Further research needed stem cells (AFSCs) - No ethical problems (latest discovery) - Possibility of preservation as lifelong autologous stem cells together with other perinatal stem cells - Possibility of ante-natal collection by amniocentesis or chorionic villous sampling

Ideally, a stem cell for RM applications should fulfill the following criteria: - Be available in abundant quantities; - Able to be harvested from a suitable donor using a minimally-invasive procedure; - Able to be differentiated along multiple cell lineage pathways in a regulatable manner; - Able to be transplanted in either an autologous or allogeneic manner with safe and efficacious outcomes; and - Able to be manufactured in accordance with current GMP Guidelines [7]. Furthermore, stem cells can be categorized based on their differentiation potential, as shown in Table 1.4.

Table 1.4 Categorization of stem cells [36] Potency Cell forming abilities Example Totipotency All types of cells Zygote Pluripotency All types of cells that arises from the three germ layers Embryonic stem cell Multipotency All cells from a single germ layer Mesenchymal stem cell Oligopotency Few types of cells that belong to their residing organ Lymphoid cell or tissue Unipotency One specific type of cell Skin cell

Embryonic stem cells (ESCs)

The most primitive stem cells derive from the descendants of the zygote, a totipotent stem cell (able to give rise to all cell and tissue types). These cells form the embryo. Four days after fertilization, these cells begin to specialize, forming a hollow ball of cells called the inner cell mass (ICM). ICM cells are pluripotent because they can differentiate into almost all cells that arise from the three germ layers (ectoderm, mesoderm and endoderm). Human embryonic stem cells (ESCs) derive from the ICM, and they are also pluripotent [37].

Somatic stem cells

Most tissues have multipotential stem cells capable of producing a limited range of differentiated cell lineages. For example, the central nervous system has stem cells that have tri-lineage potential (capable of generating neurons, oligodendrocytes, and astrocytes). These multipotent localized stem cells are called somatic stem cells (or adult stem cells). They have the capability to renew themselves and differentiate into specialized cell types of 10

the tissue they originate from. Multipotent adult stem cells, for example mesenchymal stem cells (MSCs) are more limited in descending lines and only differentiate to cells from a single germ layer: the mesoderm layer [38].

Mesenchymal stem cells (MSCs)

It is generally agreed that in an embryo, an MSC is a multipotent progenitor cell whose progeny eventually gives rise to skeletal tissues: bone, cartilage, ligament, tendon, marrow stroma and connective tissues [39]. Their progeny is affected by several factors, both intrinsic and extrinsic, that control the molecular and cellular patterns of expression that results in specific tissues that perform specific functions based on their molecular repertoire. The middle embryonic layer, the mesoderm, gives rise to all the body’s skeletal elements. These cells have the ability to spread and migrate during early embryonic development between the ectodermal and endodermal layers. This ability is the key element of all wound repair in adult organisms involving MSC in skin, bone or muscle [39]. MSCs are located in most tissues and can be isolated from various adult and embryonic tissues. The most commonly studied ones come from bone marrow and adipose tissue [40].

MSCs tri-lineage differentiation

MSCs can be differentiated into different cells from mesodermal lineage, such as osteocytes, adipocytes, and chondrocytes (Figure 1.2).

Figure 1.1 Differentiational ability of mesenchymal stem cells [41]

Differentiation in vitro depends on the culture conditions. Introduction of growth factors, such as from TGF-β (transforming growth factor beta), result in chondrogenic differentiation, while dexamethasone, insulin, isobutyl methyl xanthine, and indomethacin will cause adipogenic differentiation. To induce osteogenic differentiation, dexamethasone, ascorbic acid, and β-glycerol phosphate must be present in the culture media [42]. The importance of these constituents is becoming increasingly clear for in vitro cultures, but the in vivo biochemical environment has not been well characterized and the in vivo driving force for osteogenic differentiation remains unknown [42].

Induced pluripotent stem cells

The cloning of Dolly the sheep demonstrated that nuclei from mammalian differentiated cells can be reprogrammed to an undifferentiated state by trans-acting factors present in the oocyte (egg cell) [43]. Differentiated cells can be reprogrammed to a pluripotent state by transferring nuclear content into oocytes or by fusion with ESCs. Takahashi and Yamanaka found that introducing four factors, Oct3/4, Sox2, c-Myc and Klf4, causes a somatic stem cell to exhibit the morphology and growth properties of ESCs and to express ESC marker genes [44]. These induced cell types are called induced pluripotent stem cells (iPSC). The use of ESCs causes destruction or manipulation of an embryo. Therefore, much controversy has surrounded their use. Since iPSCs can be derived from adult tissues, the need for embryos can be bypassed and it is possible that each individual could have his or her own pluripotent stem cell line. This technology has not yet seen therapeutic use but has much potential.

2D vs 3D cultures

Cell cultures in two dimensions (2D) have been undertaken in thousands of laboratories worldwide over the past four decades. However, these cultures are primitive and do not reproduce the anatomy or physiology of any in vivo tissue. Entering the third dimension (3D), is clearly more relevant, but increases the level of complexity. The transition from 2D to 3D cultures requires a multidisciplinary approach and multidisciplinary expertise [45]. In conventional monolayer culture (2D), the transport of bioactive molecules is different from that in a 3D system [46], and thus, the patterning is biased. In a 2D culture, synthesized messenger substances can be transported in the liquid (for example media or plasma). In 12

such complex liquid environments, the diffusion coefficient of proteins, urea, creatinine, and uric acid is around 30% lower than for water at 37°C [47]. When looking at components of the ECM, there is another level of complexity due to additional mechanisms such as protein binding. When one protein, VEGF-A, is in the presence of a binding site (which is provided by the ECM structure), it contributes to vessel branching. Vessel branching is essential for formation of a dense vascular network. This results in 2D cultures being far too artificial and lacking many parameters known to be important for accurate tissue structure and physiology [47].

Bioreactors

TE uses sophisticated bioreactor systems to control process parameters, and, hence, provide optimal culture conditions. This includes several single-process steps (Figure 1.3). First, cells from biopsy are isolated, then they are expended in vitro, a 3D construct (scaffold) is generated and matured, and, finally , the construct is used as a test system or graft. A bioreactor system is a well-known device in modern TE and plays a vital role in tissue development. Bioreactors provide an in vitro environment that has physiologically tissue- specific properties for cells to mature. With bioreactor technology, various functional tissues of different types can be generated and cultured in vitro [47]. During cell expansion, bioreactors have the potential to decrease costs and improve process efficiency by fulfilling labor-intensive manual tasks. Additionally, controlled dynamic cell culture conditions overcome major drawbacks of static 2D cell culture conditions [47]. For example, the continuous perfusion of a culture system during cell expansion in 3D results in improved consistency, in terms of nutrient and waste-product distribution [48]. Areas with a high concentration of metabolic products in the proximity of the cells can be actively disrupted and supplied with fresh cell culture media. In comparison to a manual cell culture process, the molecular composition of the media does not vary due to the periodic exchange of the cell culture media. With the beginning of the tissue generation phase (phase IV in Figure 1.3), the culture conditions are switched from 2D to 3D, and at that point TE starts and conventional cell culture ends. As mentioned in the discussion of 2D vs 3D scaffolds, in vitro tissue development inherently relies on the principles of spatial and temporal organization [49].

Properties of bioreactors

Bioreactors are usually used to facilitate, monitor, and control biological or biochemical processes. [50]. The parameters that determine growth include temperature, oxygen concentration, pH, nutrient concentration and biochemical- and mechanical stimuli. Closed system bioreactors offer greater advantages with respect to good manufacturing practice (GMP). The inner space is usually adapted to spatial surroundings of incubators to provide favorable conditions for cultures, for example 99 % humidity, 37 °C and 5 % CO2 [17]. Those parameters can be monitored and changed over a defined operating range, as desired by the user. 14

Figure 1.2 Paradigm of TE: I: Biopsy. II: Isolation of cells. III: Expansion of cells using standard cell culture methods. IV: Cells seeded onto scaffolds. V: Maturation phase. VI: a) Tissue can be used as an implant b) or as a test system. A bioreactor system can be used in stages III and IV [47]

One crucial aspect that needs to be addressed during design is the diffusion limit. This is not a problem in monolayer cultures, but must be addressed in multilayered cultures. Bioreactors should provide all cells with nutrients, oxygen, and biophysical stimulus. The supply of nutrients becomes critically limiting for the culture of tissues that are thicker than 100 to 200 µm [51] (one MSC is around 25 µm in diameter, on average [52]). This is mainly a problem

15 in static cultures in bioreactors. Bioreactor systems based on fluid flows have been shown to be beneficial for cell seeding compared to static seeding methods [53], [54]. Shear stress generated by turbulent flow or perfusion stimulates proliferation and differentiation of human osteoblasts by activation of extracellular signal-regulated kinases-dependent and other pathways [55]. Some types of bioreactors are used to enhance osteogenic cellular differentiation by simulating biophysical forces that occur in vivo. The complex stimuli occurring in vivo can be hard to mimic, but can be somewhat quantified. Forces applied to bone during body movement results in changes of hydrostatic pressure, direct cell strain, fluid flow-induced shear stress and electric fields [56]. Bone cells are more sensitive to mechanical deformation than most other cells [57]. In general, mechanical loading increases bone formation and bone mass.

Bioreactor materials

Materials that are in contact with cells or are near cell cultures must be certified as biocompatible. The most common materials are synthetic polymers, for example polymethylmethacrylate, polyoxymethylene or polysulfone that are able to withstand sterilization techniques [17]. The housing itself can be any material, since it is not in contact with living material, metals are often used because of their strength.

Bioreactor systems in tissue engineering

Many types of bioreactors are used for TE purposes. This section summarizes the most common types.

Systems using hydrodynamic shear stress

Hydrodynamic shear stress bioreactor systems rely on the shear stress on scaffolds resulting from fluid flow. Local internal shear stress in scaffolds is not only influenced by medium flow rate, but also by porosity, dimensions and geometry of the scaffold, construction material, degree of interconnectivity, and the viscosity of the medium [17].

1.4.2.1.1 Spinner flask bioreactor

The spinner flask bioreactor is the simplest and cheapest bioreactor system design (Figure 1.4).

Figure 1.3 The spinner flask bioreactor [17]

Scaffolds are placed in the middle of the flask and held in place by a needle connected to the lid of the flask. The medium is stirred either by an internal rotating plate or by a magnetic stirrer. This system requires an external incubator to control temperature and oxygen/carbon- dioxide levels. Two angled side arms guarantee oxygenation of the medium. The shear stress is dependent on stirring speed. A stirring speed of 50 rpm was used in a study with collagen and silk scaffolds seeded with human MSCs [58]. One drawback of spinning systems is the possibility of formation of a dense superficial cell layer, which may hinder oxygen and nutrient supplies to the scaffolds, resulting in a lack of oxygen and nutrients reaching the middle of the scaffolds. Another drawback is the dishomogeneous shear stress on the scaffolds [17].

1.4.2.1.2 Rotating bioreactor systems

The shear stresses generated by laminar flow along the horizontal axis of a rotating vessel provide a sufficient increase in nutrient and waste diffusion through layers of cells. Different designs of rotating bioreactor systems have been developed for dynamic 3D bone TE. Some systems have scaffolds, others have culture dishes, and others have free floating cells in the

media [59], [60]. A schematic of a rotating bioreactor system can be seen in Figure 1.5.

Figure 1.4 A rotating bioreactor system: A) scaffolds are attached to the outer vessel wall during rotation; B) rotating bed bioreactor; C) Cells are seeded on discs rotating on the horizontal axis [17]

Perfusion-based bioreactor systems:

As diffusion in rotation-based bioreactors is not well understood, perfusion bioreactors producing laminar flow were developed to enable mass transport of nutrients, waste, and oxygen through scaffolds. Perfusion systems consist of containers or chambers containing the cells or scaffolds. The culture medium is piped through tubes by a peristaltic roller motor pump, and can either flow in a closed loop or be interconnected with a reservoir medium bag or chamber. The fluid flow mode can influence effects of osteogenic stimulation. Jacobs et al. found that oscillating flow was a much less potent stimulator of bone cells than steady laminar flow [61]. Perfusion bioreactors can be categorized on whether they use direct or indirect medium perfusion (Figure 1.6).

Figure 1.5 Schemetic drawing of perfusion flow principles. A) shows indirect perfusion where the medium flows around the scaffold, while B) shows direct perfusion, where medium is forced through the scaffolds. Shear stress is therefore transferred directly to the cells within the scaffolds. Arrows represent magnitued and direction of flow [17]

In a direct perfusion system, the culture medium is forced to perfuse through the scaffold, while in an indirect perfusion system, the medium takes the path of least resistance through the chamber/container. In other words, indirect perfusion does not force the medium to go through the scaffold. Higher amounts of viable cells have been observed in the center of the constructs when using direct medium perfusion than when using indirect [62]. To complement this effect, scaffolds can be manufactured with a gradient (i.e. they can be made more porous near the surface to ease cell access and nutrient flow to center of scaffold). Evidence exists that using perfusion-based bioreactor systems in TE results in improved cellular proliferation, distribution, differentiation, and viability in scaffolds, as compared to in static cultures. However, flow rates must be chosen with care; with increasing flow rate, the shear stress increases, and at a certain threshold the cells will be washed of the surface. Thus, it is important to determine the optimal medium flow rate for each bioreactor setup with, for example, the help of computer models using finite element methods.

Mechanical strain bioreactor systems:

Bones are constantly being renewed by osteoblasts (bone-forming cells) and osteoclasts

(bone-resorbing cells). Osteocytes detect mechanical stimuli and emit signals accordingly. Although the exact mechanism behind bone growth stimuli due to mechanical load is unknown, one study showed that the number of osteocytes was reduced in microgravity environments such as in space [63]. The main idea behind mechanical strain bioreactors is to mimic the mechanical forces present during walking by using cycling compression with the same frequency as in a typical gait. In addition, they can be used to gain knowledge on how mechanical stimuli affects cells. There are many different types of mechanical strain that can be applied to scaffolds, for example bending, compressing, stretching, and twisting. Strains can be uniaxial or multiaxial, and they can either be applied in cyclic or constant manner [17]. Existing bioreactor systems for direct mechanical stimulation have shown beneficial effects on proliferation, osteogenic differentiation, and matrix formation [64]. The bioreactor system designed in this project is equipped with uniaxial compression unit and indirect laminar perfusion peristaltic flow.

Electromagnetic field bioreactors:

Electromagnetic fields (EMFs) have been used for bone regeneration in patients with osteoporosis and in limb lengthening therapies [65]. Pulse EMFs (PEMFs) has been shown to reduce bone mass loss and accelerate bone formation in vivo [17], [66]. EMFs and PEMFs arise inside the human body as a result of muscle movement. Frequencies in range of 5 to 30 Hz were observed when a test subject was standing and walking [67]. EMF affects different subcellular proliferation and differentiation-related signaling pathways and studies have shown that 15 Hz is the optimal value for osteogenic effectiveness [68]. To take advantage of EMFs, bioreactors with coils have been developed to generate electric fields. Studies have shown that EMFs induce and enhance osteogenesis in human MSCs [69].

In vivo bioreactors

Several in vivo bioreactor types have been developed to generate vascularized bone tissue using different animal models [70], [71]. Even humans can be used as “bioreactors”. For example, in an attempt to repair a lower jaw bone (mandible) of a human subject, a custom bone transplant was planted inside the latissimus muscle of an adult male patient. After seven weeks, the construct was transplanted in the lower jaw successfully. This resulted in an improvement in the patient’s quality of life [72]. To date, only a few instances of in vivo 20

bioreactors have been recorded; however, other in vivo bioreactor ideas are under development.

3D printing of bioreactor components

3D printing is a promising area for producing biomedical devices and materials. Both scaffolds and bioreactor components can be constructed with 3D printing technology. One- of-a-kind devices, implants, and scaffolds can be created. In general, a large variety of ceramic, metallic, polymeric, and composite materials can be processed using 3D printing. A key advantage is that there are no contamination issues. [30]. Chapter 1.2.2 describes 3D printing technology in more detail.

Computer drawings

It is always important to make your work look good for either publishing or advertising purposes. Drawings are also made to better understand the problem that needs to be solved. Engineers use many types of drawings for this purpose, either just a paper drawing or by using computer software. The benefit of using computer software, is more accurate representation of the real object, and it prevents measurements errors. Inventor by Autodesk can be used for this purpose.

Microcomputed tomography (µCT)

Microcomputed tomography (µCT) is a technique used to image and quantify internal structures nondestructively in 3D, from two-dimensional (2D) projections. This technique provides hierarchical biological imaging capabilities with isotropic resolution of few tens of micrometers [73]. Quantification of microstructural characteristics, such as surface-to- volume ratio, porosity, and interconnectivity, is necessary for scaffold designers [74]. Many development steps of µCT have been defined for trabecular bone studies. µCT has become the standard technique in basic research and clinical laboratories to assess 3D trabecular micro-architecture. These systems provide resolutions ranging from 5 to 100 µm, so specimens with a diameter in that range can be measured. µCT is a precise and validated technique [75-77] that has been used extensively for research projects involving bone microarchitecture and to image scaffolds intended as a bone substitution [78]. The physical principle behind µCT is based on the attenuation of X-rays, just like

21 conventional CT imaging, but on a smaller scale. Conventional CT scans capture cross- sectional images and revolve the emitter and detector around a specimen, while in µCT, the specimen being imaged usually revolves.

µCT in regenerative medicine

Attenuation in bones is largely caused by mineral crystals, so it is no surprise that scaffolds containing hydroxyapatite and calcium phosphate can be imaged very nicely. Polymers and copolymers based on polylactic acid (PLA), polyglycolic acid, polyethylene glycol and polycaprolactone have been imaged using µCT to assess their 3D microarchitecture [79]. This demonstrates how µCT techniques can be used to define the best processing settings to produce scaffolds that are targeted for replacing a specific type of trabecular bone [77]. µCT scanning of polymeric scaffolds in air is possible, but the contrast between polymeric materials and their surroundings is often too low to distinguish them from watery environments, such as culture medium. To compensate, a staining agent can be integrated into scaffolds. Images obtained from scanners are usually in DICOM (Digital Imaging and Communications in Medicine) format, which is the standard for storing and transmitting medical images [80]. A stack of 2D images can be transformed into 3D a model, using software such as Mimics Research from Materialise.

Finite elements analysis

Software and hardware technologies are being used increasingly by researchers trying to understand and visualize various physical phenomena. A finite element analysis (FEA) is a well-known method used in academia and industry for many types of analyses. It is a numerical method for solving problems in engineering and mathematical physics. For example, it can be used to analyze heat transfer, fluid flow, electromagnetic potential, and mass transport. The finite element formulation of problem, results in a system of algebraic equations for solution [81]. In the context of this thesis this work, we are most interested in compression and fluid flow analyses. When a scaffold is used in a compression bioreactor, its compressional ability of it must be considered. Scaffolds with a high elastic modulus need more effort to be compressed, therefore the compression motor of the bioreactor must be able to deliver the desired 22

compression for that specific scaffold type. The wall shear stress and hydrostatic wall pressure acting on the cells must be quantified in order to know how much can be applied. One study showed that the displacement of cells is primarily determined by cell morphology [82], and that cells bridging two struts experience 500 times higher deformation than cells only attached to one strut. For alginate, the Young’s modulus is in the range of 50 to 500 kPa, depending on the mixture of substances and manufacturing methods [83]. The density is between 1060 and 1123 kg/m3 and Poisson’s coefficient is 0.5 [84]. For collagen, the Young’s modulus can change when structures are kept in media for days; normally falls within the range of 2 to 20 kPa. Poisson’s ratio for collagen is assumed to be 0.5 [85] and the density is low, in the range of 1 to 50 kg/m3 [86], [87].

Wall shear stress

It has been established that shear stress is the main biophysical stimulus which causes cells to activate matrix production and mineralization. Mechanical stimuli derive both from compression and from the movement of fluid through the scaffold. These stimuli influence gene expression, bone tissue growth, remodeling, proliferation, and differentiation. In vivo, bone cells experience estimated shear stresses of 0.8 to 3.0 Pa during routine physical activity [82]. Shear stress values outside the physiologically relevant range might lead to a lack of osteogenic stimulation. However, much lower wall shear stress values will suffice in order to stimulate osteoblastic cells in vitro. Cartmell et al. suggested a fluid shear stress in the range of 0.05 to 25 mPa to obtain positive effects on cell viability [88]. Raimondi et al. proposed that shear stress in range of 1.5 to 13.5 mPa was required for stimulation of higher cell viability. Porter et al. reported that a shear stresses of 5 × 10-5 Pa corresponds to increased cell proliferation, whereas higher shear stress levels were associated with the upregulation of bone marker genes [89][17].

Perfusion

The perfusion rate has a strong influence on cell viability. A flow rate of 1 mL/min resulted in substantial cell death throughout scaffold using direct flow method, while lowering the flow rate led to an increasing proportion of viable cells, particularly within the interior of the construct. DNA analysis showed a significant increase in cell proliferation at a flow rate of 0.01 mL/min relative to that at 0.2 mL/min [90].

An important factor in the analysis of media flow through a chamber is the determination of whether the flow is laminar or turbulent. To determine this, a parameter called Reynold’s number (Re, -) is used. Equation 1.1 yields Reynold’s number of a liquid flowing in a pipe.

흆∙풖∙풅 푹풆 = 풉 (1.1) 흁

3 where 𝜌 [푘푔/푚 ] is the density of the liquid, 푢 [푚/푠] is the speed of the liquid, 푑ℎ [푚] is the inner diameter of the pipe, and µ [푘푔/(푚 ∙ 푠)] is the dynamic viscosity of the liquid. If 푅푒 is below 2300, the flow is laminar; if 푅푒 is higher than 4000, the flow is turbulent. If 푅푒 falls between 2300 and 4000, the liquid is in transition between laminar and turbulent flow [91].

Compression

When a scaffold is compressed, the scaffold undergoes deformation. This deformation produces interstitial fluid flow inside the scaffold, causing shear stress [92]. Compression bioreactors apply compression in cycles on scaffolds in order to mimic what occurs in vivo. The temporary deformation can be expressed by Hooke’s law (Equation 1.2) [93]:

흈 = 푬휺 (1.2)

Where 𝜎 [푁/푚2] is applied stress, 퐸 [푃푎] is the Young’s modulus of the material and 휀 [푚/푚] is the strain. Care must be taken when applying compression, because too much strain and stress can lead to structural failure or plastic deformation.

Aims and Objectives

The main aim of this project was to redesign and construct a stand-alone bioreactor system using computer aided design and 3D printing.

The main objectives of this project were to:

i. Design and construct a new prototype for a standalone bioreactor system for 3D osteogenic differentiation incorporating mechanical stimulation and perfusion

ii. Estimate and optimize shear stress from compression and fluid flow on the scaffolds inside the bioreactor, using finite elements analysis to improve further development of the bioreactor system.

iii. Make the bioreactor system user friendly by creating manual and protocols.

22 Background

This project is a continuation of the work of a former PhD student at Reykjavik University and the University of Bologna, Joseph Lovecchio. His work laid the foundation for this project [94]. He obtains his PhD degree in biomedical engineering in spring of 2019. He developed a standalone bioreactor system which was capable of performing perfusion and compression on scaffolds, maintaining constant temperature and pH values. Joseph provided the information and materials needed to continue developing the bioreactor. There were a few areas for improvement in the bioreactor system developed by Joseph Lovecchio. The main issues with it were:

Main problems with it were: • Leakage of media out of the chamber and at joints • The chamber walls bulged due to heat strain • Lack of sterility • Immobile design (hard to move materials from the bioreactor due to presence of pipes)

The work of this thesis project revolved around solving these problems.

33 Methods & Results

The first part of this thesis work involved improving the bioreactor system to the point where it was fully functional; this was necessary to complete before experiments could be run using the bioreactor. Many attempts by trial and error, along with unforeseen technical difficulties, made this first step extremely time consuming.

Bioreactor evaluation and redesign

The initial part of the evaluation process was to identify the flaws in the existing system. New versions of chambers were drawn and designed Autodesk Inventor. Autodesk Meshmixer software was also used, because of its ability to manipulate .stl files, which are the file types sent to the 3D printer for printing. When the 3D drawing was ready, it was exported as .stl file for printing. It is not possible to manipulate the .stl file in Inventor. Meshmixer was therefore used when changes had to be made to the .stl file.

3D printing

Full access to a 3D printer at Reykjavik University (model: Stratasys objet260 connex3) provided the possibility and freedom to keep working on changes to the bioreactor chamber. A variety of materials were available for printing. The 3D printing software included with the printer is simple and efficient. A simulation of the printer tray appears on the software interface which makes it easy to import models and to rotate and scale them as desired. Various designs of the chamber were made. All versions of the 3D-printed chambers had to be properly pretreated before they could be used with cells. A protocol was written on how to treat the chamber using ultraviolet (UV) light, a waterjet and a salt bath. This protocol can be found in Appendix B.

Screw mechanism addition

The main problem with the existing chamber was the media leakage. This was solved via

the addition of a screw mechanism on the port holes of the chamber. Figure 3.1 (left) shows the Luer lock style (a standardized system for fluid fittings ports) screw port scanned with a µCT scanner, along with the computer-generated mesh (right side of Figure 3.1). The mesh was edited to fit into the chamber using Autodesk Meshmixer, as seen in Figure 3.2. With the addition of the screw mechanism, the chamber was more mobile. It is sterile when every screw port is sealed with a Spiros connector, and if pH and temperature sensors are in place.

Figure 3.1 Left: A Luer lock style connector. Right: A computer generated image of the connector on the left. This was made with a µCT scanner and Autodesk Inventor software

Figure 3.2 Screenshot from Autodesk Meshmixer when scew ports (grey, lime, purple, green and cyan colored) were inserted onto the chamber

Numerous trial and error attempts were made prior to finalizing the screw connector mechanisms in order to find the right combination of materials and wall thickness., and 3D printed to find right combination of materials and wall thickness. This will be discussed further in later subchapters. 28

Spiros and the Autoclave connectors

The screw mechanism on the chamber was designed to fit Spiros® (blue) and Microclave® (white) connectors from ICU Medical (www.icumed.com) (Figure 3.3). When the connectors are not connected, they are closed and sealed. The connectors are in an open state only if a female connector is attached to the Microclave or if a male connector is connected to the Spiros. Both of these connectors have been validated as microbiologically-closed devices.

Figure 3.3 Left: A Spiros® connector. Right: A Microclave® connector. They are both made by ICU Medical

Screw port tests

On the first versions of the 3D printed chamber, the screw mechanism proved to be extremely brittle. To solve this problem, a strand containing nine screw ports, each different from the other, either in wall thickness, material mixture, or both, were created and 3D printed. In most cases, it was the inner piece of the screw port that broke, but tests were made with many combinations of both thicker and thinner walls and thicker and thinner inner pieces. To complicate matters even more, the brittleness of the material changed whether the material had been exposed to UV light (part of cleaning protocol) and depending on which direction screw ports were facing during printing. Three different strands were made, with each new one narrowing down to best composition. The cleaning protocol (Appendix B) was followed prior to testing. Testing was performed manually. The final choice of screw connector had to be both strong and not too brittle. Table 3.1 lists up all combinations tested. One of the strands can be seen in Figure 3.4 which shows how the leakage/strength test was done.

Figure 3.4 Screw test strand (left). Testing the screw ports for strength and leakage (right)

Table 3.1 Three 3D-printed strands with nine different ports for determining the best material and thickness of screw mechanism. Materials used in test the strands were: MED610, Veroclear hard plastic, and Agilus 30 rubber. Some pieces were also coated with MED610 (adds 0.3 mm layer on the surface). The best combination is shown in bold.

1ST 3D-PRINTED STRAND 2ND 3D-PRINTED STRAND 3RD 3D-PRINTED STRAND PRINTED FACEUP PRINTED FACEUP PRINTED SIDEWAYS

1 Original (for comparison) 1 Original (for comparison) 1 Original (for comparison)

2 Only thicker walls, 100 % 2 Only thicker walls, 100 % 2 Only thicker walls, 100 % MED610 MED610 MED610

3 Thicker walls and thick 3 Thicker walls and thick 3 Thicker walls and thick middle. middle. 100 % MED610 middle, 100% MED610 100 % MED610

4 Thicker walls and thin 4 Thicker walls and thin middle. 4 Thicker walls and thick middle. middle. 100 % rubber 75 % Veroclear, 15 % rubber 100 % rubber + coated 5 Thicker walls and thin 5 Thicker walls and thin middle. 5 Thicker walls and thick middle. middle. 100 % Veroclear 100 % Veroclear 85 % Veroclear, 15 % rubber

6 Thicker walls and thick 6 Thicker walls and thick 6 Thicker walls and thick middle. 100 % rubber middle. 70 % Veroclear middle. 85 % Veroclear, 15 % rubber + coated 7 Thicker walls and thick 7 Thicker walls and thick 7 Thicker walls and thick middle. middle, 50 % Veroclear, 50 middle. 85 % Veroclear, 15 % 95 % Veroclear, 5 % rubber % rubber rubber 8 Thicker walls and thick 8 Thicker walls and thick 8 Thicker walls and thick middle. middle, 70 % Veroclear, 30 middle. 95% Veroclear, 5 % 95 % Veroclear, 5 % rubber + % rubber rubber coated 9 Thicker walls and thick 9 Thicker walls and thick 9 Thicker walls and thick middle. middle. 100 % Veroclear middle. 100 % Veroclear 100 % Veroclear + coated

The best screw port proved to be one with thicker walls and thick middle, 85% Veroclear hard plastic, 15% Agilus30 rubber, and coated with MED210 (shown in bold in Table 3.1).

Chamber version development

As mentioned before, bulging of the media chamber was a major problem in the previous bioreactor design. We therefore had to redevelop the chamber using a trial and error method. A total of nine chambers were designed and 3D printed, each one having either minor or major improvements. These different printouts have been narrowed down to four versions of the chamber for simplification. The following sections provide an overview of the chamber versions, including what was new and what went wrong. All changes were made in Inventor and Meshmixer. Figure 3.5 (right chamber) shows the chamber before improvements were made, compared to a later version (left). The lack of a screw mechanism in earlier versions resulted in isolation problem (leakage along joint of rubber ring and connection pipes).

Figure 3.5 Comparison between chambers before starting redesigning (right) and after improvements were made (left)

Version 1

This was the first version which implemented a screw mechanism. The screw ports proved to be very brittle and too thin, and they were too close together, preventing Spiros (blue connectors) from connecting without touching each other (Figure 3.6).

Figure 3.6 Version 1; connectors too close together

Version 2

An increased distance between the top screw ports provided enough space in between ports. An increased wall thickness of 1 mm solved the wall bending problem. Screw ports proved to be still too brittle. Figure 3.7 shows the fully cleaned version 1 (left), and version 2 straight out of the printer (prior to cleaning) (right).

Figure 3.7 Version 1 chamber (left) fully cleaned and version 2 chamber (right) straight out of the printer. The yellow tint is caused by UV radiation and can be reduced significantly by using light treatment (Appendix B) 32

Version 3

In version 3, a pipe system connected to the two far right ports on top, leading down to bottom of chamber (Figure 3.8) was implemented. In this version, the best screw port design was used (see Section 3.1.2 – Screw port tests).

Figure 3.8 Addition of pipe system connecting two far right screw ports down to the culture medium

Version 4

Version 4 incorporated the addition of an internal plate system (Figure 3.9) on the bottom of the chamber. And other minor improvements (including narrowing dimensions to make sure the plates fit securely).

Figure 3.9 The internal plate (left) and the supporting plate (right). Further component description can be found in Appendix A – components overview

Implementing an internal plate system had the following benefits: - There was no need to print a whole chamber to accommodate different sizes of scaffolds, only a new internal plate. - Placement of scaffolds can be done outside of the chamber itself. - Molding of hydrogel scaffolds is possible. Figure 3.10 shows chamber without and with the internal plates in place.

Figure 3.10 Empty chamber (left) and chamber with the internal plate system (right)

Figure 3.11 shows the configuration of the internal plate and support plate when molding hydrogel scaffolds.

Figure 3.11 Internal plate and supporting plates in a hydrogel molding configuration. The molding takes place in between teeth near the bottom

Final chamber configuration

The chamber has several inputs and outputs and other openings with different purposes. A diagram of chamber inputs and outputs can be seen in Figure 3.12. There are six openings which require a screw mechanism connector, plus two others for temperature and pH sensors. Additionally, there are three openings on the back of the chamber: one for the big screw of the motor and two for screws holding the motor in place.

Figure 3.12 A schematic view of the port hole purposes of the bioreactor chamber.

Professional drawings using Inventor

Professional drawings of the chamber were made in Inventor. First drawing of the chamber without the internal plates was generated and can be seen in figure 3.13. All dimensions needed to replicate the design are included. Next, a drawing was made, of the chamber with the internal plates in place (figure 3.14). No dimensions are needed here because they are identical to the first sketch. Finally, drawings were made of individual plates of the chamber (figure 3.15).

chamber made in Inventor. Dimensions are in Dimensions are inInventor. millimeters. made chamber

Schematic Schematic drawingbioreactor the of

3 Figure Figure

Schematic view of the bioreactor chamber including the internal plates. including view Schematic chamber internal the of bioreactor the

3 Figure Figure

Schematic view of other chamber components. Dimensions are in millimeters. are Dimensions view Schematic components. chamber of other

3 Figure Figure

Finite element analysis

In this project, a finite element (FE) approach was used to analyze several phenomena. As discussed in earlier subsections, 47.5B bioactive glass scaffolds were used in the bioreactor and, therefore, the FE analysis was made with that scaffold type in mind, but also other types such as collagen and alginate. COMSOL Multiphysics software (Comsol Inc.) was used for this analysis. This program allows the user to apply many physical stimuli simultaneously.

Meshing and preprocessing

The first step in the FE analysis was to create the geometry of the space that the liquid media occupies inside the bioreactor chamber. Figure 3.16 shows the geometry created.

Figure 3.16 Geometry representing liquid media space inside the bioreactor chamber

The density of the media was considered to be 1007 kg/m3 (similar to the density of water, which is 1000 kg/m3), also, the viscosity was comparable to water [95]. Since water is a premade element in the software, so water was assigned as the media space. The next step was to mesh the geometry. The software offers six different types of meshing, from extra coarse to extra fine. Working with extra coarse mesh takes the least computational time, but yields the most inaccurate results, and vice versa for the extra fine mesh. Computing the extra fine mesh processed for 1,5 hours without reaching the end of the run, so the run was killed. It was decided that a compromise had to be made between computational speed and accuracy. Figure 3.17 shows four different element quality histograms.

Figure 3.17 Histograms for differently detailed meshes of the media geometry

The overall mesh quality is better when the histogram is skewed more to the right side of the graph. There was not a significant difference seen between the histograms of the fine mesh and the normal mesh, so for this analysis a normal mesh size was used. The final mesh can be seen in Figure 3.18.

Figure 3.18 Computer generated mesh of the media area of the chamber

Flow inputs can be assigned to the surface in COMSOL, along with their desired velocity. Equation 3.1 was used to convert the volumetric flow rate Q to velocity v (knowing the area A).

푸 [풎ퟑ/풔] ퟒ푸 풗 [풎⁄풔] = = (3.1) 푨 [풎ퟐ] 흅풅ퟐ

With 푄 = 2 푚퐿/푚푖푛 = 0.033 푚퐿/푠 = 3.33푒 − 8 푚3/푠 and media entrance diameter of 푑 = 2.5 푚푚, the velocity is 푣 = 6.78푒 − 3 푚/푠. With that in mind, the Reynold’s number was used to determine whether the input flow is laminar or turbulent. Equation 1.1 yields Reynold’s number.

With a density ρ = 1000 kg/m, velocity v = 6.78 mm/s, diameter of pipe dh = 2.5 mm and dynamic viscosity µ37°C = 0.679 mPa·s [96]. Reynold’s number is close to 0. If Reynold’s 40

number is less than 2300, the flow is laminar. Simulations were carried out using the above- determined parameters.

Compression

To achieve 1 % uniaxial deformation of the scaffolds, the Young’s modulus of the material must be known. By using the definition of the Young’s modulus (Equation 3.2) the required compression can be determined.

풔풕풓풆풔풔 흈 푬 = = => 흈 = ퟎ. ퟎퟏ푬 (3.2) 풔풕풓풂풊풏 ퟏ%

Compression was applied on the scaffolds (the 4 squares seen in figure 3.16). The average shear stress within the scaffold volume was evaluated. Shear stress was evaluated for scenarios of only perfusion, only compression, and both combined. A few types of scaffolds were evaluated, including alginate, and collagen.

Parameters

Young’s modulus, Poisson’s coefficient and density were estimated as a function of porosity for the 47B glass. Table 3.2 shows the relevant equations as a function of porosity, as well as the constants for bulk glass with no pores.

Table 3.2 Relevant equations as a function of porosity for mechanical parameters

Parameter Parameter as a function of porosity equation Bulk parameter value

2 Young’s 퐸 = 퐸0(1 − 푝) [97] 퐸0 = 72.9 퐺푃푎 modulus (푬) Poisson’s 2 1.21 휈 = 0.2679 (1−푥3) 0 coefficient ν = 0.5 − (3−5∗푥)∗(1−푥) 1−푥 4∗(((1−푠)∗ )+푠∗( )) (흂) 2∗(3−5∗푥)∗(1−2∗푣0)+3.∗푥∗(1+푣0) 3∗(1−푣0) [98] 3 Density (흆) 𝜌 = −𝜌0푝 + 𝜌0 𝜌0 = 2640 푘푔/푚

Figure 3.19 shows these three parameters as a function of porosity.

Figure 3.19 Young’s modulus, Poisson’s ratio and density as a function of porosity. The y coordinates seen is the graphs, are the obtained values for the bioactive glass scaffolds used. The Matlab code can be found in Appendix D

The average porosity in all six scaffolds was measured to be 38.52 %. The Young’s modulus, Poisson’s coefficient and that density was calculated to be: 퐸 = 27.13 퐺푃푎, 푣 = 0.126 and 𝜌 = 1610 푘푔/푚3, respectively. These parameters were used throughout the FE analysis.

Shear stress

The volume averaged shear stress for five scaffold types tested, under conditions of compression alone, perfusion alone and both compression and perfusion, can be seen in table 3.3.

Table 3.3 Results from the FE analysis. Shear stress evaluation from compression, perfusion, and both combined.

Volume average shear stress during 1% compression from… Scaffold …Compression …Perfusion …Both Perfusion/Compression [Pa] [Pa] combined [Pa] shear stress 47.5B Bioactive glass 6000 0.261 6000 0.00% 47.5B with lowered Young's modulus 1 60 0.261 60.3 0.43% 47.5B with lowered Young's modulus 2 0.6 0.261 0.861 43.44% Alginate 0.144 0.261 0.405 180.77% Collagen 0.00577 0.261 0.266 4519.11%

Flow analysis

The velocity of media flow was analyzed both in a plane near the entry point and on a plane on the scaffold level. Experiments with two other types of shapes were carried out. The first shape represents the current chamber system (Figure 3.20 - a). The second shape included tapered edges (Figure 3.20 - b), and the third included curved edges (Figure 3.20 - c).

Figure 3.20 The shapes of chamber used for FE analysis, a) Original chamber, b) chamber with tapered edges, c) chamber with rounded edges

Figure 3.21 shows total flow velocity in all chambers in a horizonal plane at the inflow level.

Figure 3.21 Total velocity flow in meters per seconds in all directions at inflow level

Figure 3.22 shows the total flow velocity in all chambers in a horizonal plane at the scaffold level.

Figure 3.22 Total velocity flow in meters per second in all directions, at scaffold level

Figure 3.23 shows flow in the x-direction (inflow stream direction) in a horizonal plane at the inflow level.

Figure 3.23 Total velocity flow in meters per seconds only in x-directions, at inflow level 44

Figure 3.24 shows flow in the x-direction in a horizonal plane at the scaffold level.

Figure 3.24 Total velocity flow in meters per seconds only in x-directions, at scaffold level

44 Discussion

Tissue engineering has come a long way since its initiation. The thought of being able to replace a damaged or infected section of bone with engineered tissue made from a patient’s own stem cells is not far from becoming a reality. All steps, from the biopsy to the implantation of the engineered tissue, must be clear and well-thought-out processes. Evidence exists that the use of perfusion-based bioreactor systems in bone TE results in improved cellular proliferation, distribution, differentiation, and viability in the interior of scaffolds when compared to static cultures [17]. Mechanical compression has been found to increase differentiation and proliferation [99]. Because of technical difficulties, both with the bioreactor itself and the µCT scanner, the biological part could not be completed with satisfactory results. Therefore, a description and discussion of test runs as well as static 2D culture can be found in appendix C. Our bioreactor system has proven to have potential for above-mentioned applications. Although some technical issues still remain, major improvements have been made to the system in terms of isolation, functionality and user accessibility through manual and protocols.

Bioreactor

The aim of this project was to continue to improve upon the design of the bioreactor and to make the bioreactor system operational. In this section, potential future improvements and general ideas about the system are discussed.

Bioreactor status

The bioreactor system is very close to become operational. It was only due to the lack of time for me to run a successful run in the bioreactor. The short-term plan should be to order new connectors (both pipes and Spiros/Microclave) from ICU Medical, through their catalogue, since existing pipes/connectors are worn out from frequent use. Next subsection brings up ideas for a long-term plan for the bioreactor system. 46

Future improvements

Several chamber improvements are possible, as summarized in the following list.

- The design of the box could be changed to have a more streamlined appearance, with the corners being either tapered or rounded. The reason for this suggested change is that it is desired to have all of the media in perfusion. Having sharp corners can cause some parts of the media to be left out of the perfusion as will be discussed in the finite element discussion (Section 4.2).

- The number of unnecessary ports could be reduced to decrease the risk of contamination.

- More research should be considered for the chamber material. It would be ideal to have it transparent, which could be obtained by using other material than MED610 or to keep experimenting how to make it more transparent.

For the pH sensor port: - Implementation of an attached pH sensor unit to the chamber could be done in such a way that the sensor can be detached from the chamber without leaving it open. Some possible ways to make this happen: o Implementation of a rail guiding system that allows the sensor to slide down into the chamber, along with a valve (out of rubber or more advanced system) that keeps the inner environment closed off from the surroundings when sensor is not in place, but opens when sensor is pushed down into the chamber. o Use of an internal pH meter attached to the inner surface of the lid, which would only require a screw-on mechanism on the outside of the lid for connecting to the sensor. This idea is more complicated than the previous one mentioned as it requires a specially-designed pH meter. - Or inclusion of a wireless pH sensor inside the chamber: This would eliminate all problems with removing the sensor. Such devices already exist, for example the wireless pH sensor from PASCO [100].

- Similar issues exist for the temperature sensor as for the pH sensor. A similar approach to that discussed for the pH sensor could be used to improve the design. This case is simpler, however, due to the nature of the sensor. For example, an attachment to the lid of the chamber, leading into the liquid media, could be implemented. A screw mechanism on the outside of the chamber would connect the temperature sensor from the outside of the chamber. Precautions must be taken when extending a temperature sensor, because adding extra wire can cause its internal resistance to change, resulting in measurement errors.

- Scaffold placement could be made easier by implementing a railing system for a scaffold holder that slides the scaffolds down into the chamber.

- Porter et al reported a µCT imaging-based technique to monitor 3D mineralization online over time in a bioreactor. This method can specifically assess the mechanisms of construct mineralization by quantifying the number, size, and distribution of mineralized particle formation [101]. His bioreactor was designed to fit into a CT scanner without disconnecting any tubing so that the system remained closed and sterile. This is also the case with our system, so this idea could be implemented for our bioreactor. However, some changes would be needed to make this possible. In particular, it would be ideal to have the scaffolds more accessible for the x-rays in order to achieve the most accurate results.

Other improvements: - In the current bioreactor design, there are pipes all over the place. It would be an improvement for the pipes to be placed in a more organized manner. This improvement can be made once the final positions for all system elements is established. For example, the pipes could potentially be attached to the metal case and ordered with a specifically desired length from ICU Medical.

- The current system suffers from gas inflow problems; this is one of the biggest issues with this system. Experiments were carried out in order to improve how the gas is introduced to media. In one run, gas was applied straight to the chamber. In another run, gas was applied from a reserve media bag connected to the perfusion cycle. Unfortunately, these tests were unable to be completed and further tests need to be 48

carried out. These experimental runs are described in Appendix C.

- The bioreactor is currently powered by a desktop computer. This is a problem because when electricity outages occur, the system shuts down permanently. Switching from a desktop to a laptop with battery backup will eliminate this problem.

Existing bioreactors

There are many bioreactor systems commercially available that generate mineralized cell constructs (Table 4.1). The cheapest systems are, as mentioned in the introduction, the spinner flasks. The Bell-Flow™ spinner flask from Bellco Biotechnology is manufactured from borosilicate glass and is available in volumes from 100 mL up to enormous 36 L containers [102]. The containers ranging from 100 mL to 1 L are more common and these containers include internal bearing assemblies.

Table 4.1 Selected commercially available bioreactor systems [17]

Company Product Type (features, options) Bellco Biotechnology (Vineland, NJ) Bell-Flow™ Spinner flask Corning® Lifesciences (Lowell, MA) ProCulture® glass spinner Spinner flask flask (disposable or autoclavable) MINUCELLS and MINUTISSUE Tissue engineering Perfusion bioreactor Vertriebs GmbH (Bad Abbach, Germany) container (indirect perfusion) MINUSHEET® tissue carrier BISS (Bangalore, India) OsteoGen Bioreactor Perfusion bioreactor (pulsatile fluid flow stimulator, µCT) Zellwerk GmbH (Oberkrämer, Germany) RP bioreactor Rotating bed bioreactor (GMP conform)

Synthecon, Inc. (Houston) RCCMAX Bioreactor Rotating wall vessel Flexcell International Corporation e.g., BioPress™ Systems using tension, (Hillsborough, NC) Compression compression and shear Plates, FX-5000™ stress Compression System

Corning® Lifesciences offers disposable spinner flasks. This system comes ready-to-use with a paddle and integrated magnet, reducing the need for time-consuming assembly, cleaning or reassembly. These flasks don’t have built-in internal assembly, which makes it impossible to be used for scaffolds [103].

MINUCELLS and MINUTISSUE GmbH offers an indirect perfusion system that requires fixed shape of biomaterial scaffold used. Different types of scaffolds can be used simultaneously in this system [104]. The OsteoGen Bioreactor from BISS is a direct perfusion system. This bioreactor includes a small chamber which fits only for a cylindrical scaffold, and it incorporates x-ray window for µCT mineral quantification [105]. The bioreactor system from Zellwerk GmbH is the RP bioreactor. This system has perfusion mode and a rotary bed which is set by a magnetic drive. By increasing the rotation of the system, cell can be detached from the scaffolds on purpose for harvesting purposes. The complete cultivation system comprises a bioreactor, GMP breeder and a control unit [106]. The rotating bioreactor system from Synthecon Inc were originally designed by the National Aeronautics and Space Administration (NASA). They were designed to be either autoclavable or disposable units. Synthecon also makes custom designed bioreactors, tailored to customer needs. Systems using compression or tension are available from, for example, Flexcell International Corporation [107]. The FX-5000™ systems applies cyclic or static strain to cells cultured on flexible-bottomed culture plates. Special devices allow to observe signaling responses upon strain stimulation in real-time on a microscope stage [108][17]. These existing solutions can all be used for osteogenic differentiation evaluation purposes. The main difference between them, is the complexity level. The general rule is: the more complex the system is, the more the manufacturer is trying to mimic conditions inside the body, and/or the manufacturer is making the system more automatic. Our bioreactor system can be categorized with the more complex systems, since it has both compression, perfusion, and it has optimal gas composition and temperature.

Finite element analysis

According to previous studies, the shear stress value in the scaffolds should be in the range of 5 µPa to 25 mPa. Perfusion alone with a 2 mL/min flow rate provides a volume averaged shear stress of 0.261 Pa, which is much higher value than the desired maximum. This results in cells washing off the surface of the scaffold. Considering the shear stress from compression alone, the average volume shear stress required is too large for the hard materials (47.5B bioactive glass variations) to handle. However, for the collagen scaffold, 5.77 mPa is required to reach 1 % compression, which is within the desired range. So, for this particular bioreactor system, scaffolds with a Young’s modulus in the same range as collagen are recommended for bone cell culture growth. Note that the simulation approximated the scaffold as a bulk material, while still using the elastic properties of the porous material. This results in inaccurate values for shear stress of the scaffolds. For more accurate results, it would be better to, for example, use µCT scans of the scaffold and thereby obtain accurate tomography of the scaffold. Then computer technology could be used to evaluate the shear stress internally inside the porous scaffold. Figures 3.21 and 3.22 show how fast the media travels inside the chamber during perfusion. The liquid enters from left inflow pipe and exits at same rate at the right outflow pipe. Figure 3.21 shows the flow speed in every direction at the inflow level, while Figure 3.22 shows the flow speed at the scaffold level. There is a clear difference in the speed. The reddest part shows speed up to 0.2 mm/s, with a high-speed large zone near the inflow. The bluest part shows speeds from 0-0.02 mm/s, which is slow but still moving. The do not appear to be any significant differences between the three chamber shapes, except that the curved and tapered designs don’t seem to have as slow speeds in the corners as the original one. Figures 3.23 and 3.24 show only the flow speed in the x direction of the chamber. It is interesting to see how the media seems to spiral around the inflow (blue is backwards flow) in all designs. This can cause a lack of mixing of the chamber liquid, resulting in some parts of the chamber not receiving the gas exchange needed for cell viability. Further design improvements should be considered to overcome this problem. The three different chamber shapes do not show much difference, other than perhaps less mixing in the corners of the original chamber shape. Future designs of the chamber should therefore include more tapered or rounded corners.

Conclusion

This project was successful as a development of the bioreactor system. It was time- consuming task and a lot of time went into experimenting and developing the bioreactor system. Further development needs to take place, whether it is with this design, or a new approach. The author has gained a lot of experience as well as knowledge in the TE field working on this project, for example in: - Cell handling: o Seeding of MSCs o Counting cells o Harvesting cells o Performing osteogenic differentiation on MSCs - Working with medical imaging software (Materialise Mimics): o Analyzing µCT scans using histograms - Performing FE analysis - 3D printing: o General maintenance of 3D printer o Printing biomaterials o Cleaning of printed materials - Computer designing using Inventor o General designing o Professional drawings This project has provided knowledge to continue developing as a biomedical engineer for me personally.

55 Bibliography

[1] J. P. Vacanti and C. A. Vacanti, Principles of Tissue Engineering. Boston: Academic Press, 2014.

[2] L. Greenemeier, “Blade Runners: Do High-Tech Prostheses Give Runners an Unfair Advantage?,” Scientific American. [Online]. Available: https://www.scientificamerican.com/article/blade-runners- do-high-tech-prostheses-give-runners-an-unfair-advantage/. [Accessed: 01-Oct-2018].

[3] C. Mason and P. Dunnill, “A brief definition of regenerative medicine,” Regenerative Medicine, vol. 3, no. 1, pp. 1–5, Dec. 2007.

[4] “Dr. William Haseltine on Regenerative Medicine, Aging and Human Immortality.” [Online]. Available: https://www.lifeextension.com/Magazine/2002/7/report_haseltine/Page-02?p=1. [Accessed: 01-Oct-2018].

[5] A. S. Colwell, M. T. Longaker, and H. P. Lorenz, “Mammalian Fetal Organ Regeneration,” in Regenerative Medicine I: Theories, Models and Methods, I. V. Yannas, Ed. Berlin, Heidelberg: Springer Berlin Heidelberg, 2005, pp. 83–100.

[6] B. V. Slaughter, S. S. Khurshid, O. Z. Fisher, A. Khademhosseini, and N. A. Peppas, “Hydrogels in Regenerative Medicine,” Advanced Materials, vol. 21, no. 32‐33, pp. 3307–3329, Sep. 2009.

[7] J. M. Gimble, A. J. Katz, and B. A. Bunnell, “Adipose-Derived Stem Cells for Regenerative Medicine,” Circulation Research, vol. 100, no. 9, pp. 1249–1260, Apr. 2007.

[8] G. Sampogna, S. Y. Guraya, and A. Forgione, “Regenerative medicine: Historical roots and potential strategies in modern medicine,” Journal of Microscopy and Ultrastructure, vol. 3, no. 3, pp. 101–107, Sep. 2015.

[9] J. E. Arenas-Herrera, I. K. Ko, A. Atala, and J. J. Yoo, “Decellularization for whole organ bioengineering,” Biomed Mater, vol. 8, no. 1, p. 014106, Feb. 2013.

[10] T. W. Gilbert, A. M. Stewart-Akers, and S. F. Badylak, “A quantitative method for evaluating the degradation of biologic scaffold materials,” Biomaterials, vol. 28, no. 2, pp. 147–150, Jan. 2007.

[11] A. Atala, S. B. Bauer, S. Soker, J. J. Yoo, and A. B. Retik, “Tissue-engineered autologous bladders for patients needing cystoplasty,” The Lancet, vol. 367, no. 9518, pp. 1241–1246, Apr. 2006.

[12] B. E. Uygun et al., “Organ reengineering through development of a transplantable recellularized liver graft using decellularized liver matrix,” Nature Medicine, vol. 16, no. 7, pp. 814–820, Jul. 2010.

[13] “The Difference Between Regenerative Medicine and Tissue Engineering,” 13-Aug-2018. [Online]. Available: https://www.celixir.com/the-difference-between-regenerative-medicine-and-tissue- engineering/. [Accessed: 03-Dec-2018].

[14] D. L. Butler, S. A. Goldstein, and F. Guilak, “Functional Tissue Engineering: The Role of Biomechanics,” Journal of Biomechanical Engineering, vol. 122, no. 6, p. 570, 2000.

[15] J. P. Vacanti, M. A. Morse, W. M. Saltzman, A. J. Domb, A. Perez-Atayde, and R. Langer, “Selective cell transplantation using bioabsorbable artificial polymers as matrices,” J. Pediatr. Surg., vol. 23, no. 1 Pt 2, pp. 3–9, Jan. 1988.

[16] “The Kidney Transplant Waitlist – What You Need to Know,” National Kidney Foundation, 07-Jan- 2016. [Online]. Available: https://www.kidney.org/atoz/content/transplant-waitlist. [Accessed: 24-Sep- 2018].

[17] J. Rauh, F. Milan, K.-P. Günther, and M. Stiehler, “Bioreactor Systems for Bone Tissue Engineering,” Tissue Engineering Part B: Reviews, vol. 17, no. 4, pp. 263–280, Aug. 2011.

[18] D. Logeart‐Avramoglou, F. Anagnostou, R. Bizios, and H. Petite, “Engineering bone: challenges and obstacles,” Journal of Cellular and Molecular Medicine, vol. 9, no. 1, pp. 72–84, 2005.

[19] L. Fassina, L. Visai, M. G. C. De Angelis, F. Benazzo, and G. Magenes, “Surface modification of a porous polyurethane through a culture of human osteoblasts and an electromagnetic bioreactor,” Technology and Health Care, vol. 15, no. 1, pp. 33–45, Jan. 2007.

[20] D. W. Hutmacher, “Scaffolds in tissue engineering bone and cartilage,” Biomaterials, vol. 21, no. 24, pp. 2529–2543, Dec. 2000.

[21] L. L. Hench, R. J. Splinter, W. C. Allen, and T. K. Greenlee, “Bonding mechanisms at the interface of ceramic prosthetic materials,” Journal of Biomedical Materials Research, vol. 5, no. 6, pp. 117–141, Nov. 1971.

[22] A. A. El-Rashidy, J. A. Roether, L. Harhaus, U. Kneser, and A. R. Boccaccini, “Regenerating bone with bioactive glass scaffolds: A review of in vivo studies in bone defect models,” Acta Biomater, vol. 62, pp. 1–28, 15 2017.

[23] J. van den Dolder, P. H. M. Spauwen, and J. A. Jansen, “Evaluation of Various Seeding Techniques for Culturing Osteogenic Cells on Titanium Fiber Mesh,” Tissue Engineering, vol. 9, no. 2, pp. 315– 325, Apr. 2003.

[24] M. Kruyt et al., “Analysis of the Dynamics of Bone Formation, Effect of Cell Seeding Density, and Potential of Allogeneic Cells in Cell-Based Bone Tissue Engineering in Goats,” Tissue Engineering Part A, vol. 14, no. 6, pp. 1081–1088, Jun. 2008.

[25] M. Kucharska, B. Butruk, K. Walenko, T. Brynk, and T. Ciach, “Fabrication of in-situ foamed chitosan/β-TCP scaffolds for bone tissue engineering application,” Materials Letters, vol. 85, pp. 124– 127, Oct. 2012.

[26] H. Cao and N. Kuboyama, “A biodegradable porous composite scaffold of PGA/β-TCP for bone tissue engineering,” Bone, vol. 46, no. 2, pp. 386–395, Feb. 2010.

[27] N. Sultana and M. Wang, “Fabrication of HA/PHBV composite scaffolds through the emulsion freezing/freeze-drying process and characterisation of the scaffolds,” J Mater Sci: Mater Med, vol. 19, no. 7, p. 2555, Aug. 2007.

[28] H. Yoshikawa, N. Tamai, T. Murase, and A. Myoui, “Interconnected porous hydroxyapatite ceramics for bone tissue engineering,” Journal of the Royal Society Interface, vol. 6, no. SUPPL. 3, pp. S341– S348, 2009.

[29] E. M. Sachs, J. S. Haggerty, M. J. Cima, and P. A. Williams, “Three-dimensional printing techniques - patent,” US5204055A, 20-Apr-1993.

[30] S. Bose, S. Vahabzadeh, and A. Bandyopadhyay, “Bone tissue engineering using 3D printing,” Materials Today, vol. 16, no. 12, pp. 496–504, Dec. 2013.

[31] Solid Concepts, Fused Deposition Modeling (FDM) Technology. .” [Online]. Available: https://www.youtube.com/watch?v=WHO6G67GJbM [Accessed: 10-Jan-2019].

[32] “MED610, MED620 & MED625FLX : Biocompatible 3D Printing Materials | Stratasys.” [Online]. Available: https://www.stratasys.com/materials/search/biocompatible. [Accessed: 06-Jan-2019].

[33] M. Tavassoli and W. H. Crosby, “Transplantation of marrow to extramedullary sites,” Science, vol. 161, no. 3836, pp. 54–56, Jul. 1968.

[34] E. J. Culme-Seymour, N. L. Davie, D. A. Brindley, S. Edwards-Parton, and C. Mason, “A decade of cell therapy clinical trials (2000–2010),” Regenerative Medicine, vol. 7, no. 4, pp. 455–462, Jul. 2012.

[35] T. R. Heathman, A. W. Nienow, M. J. McCall, K. Coopman, B. Kara, and C. J. Hewitt, “The translation of cell-based therapies: clinical landscape and manufacturing challenges,” Regenerative Medicine, vol. 10, no. 1, pp. 49–64, Jan. 2015.

[36] H. Montazeri, “Freeze dried platelet lysates from expired platelet concentrates support growth and differentiation of mesenchymal stem cells,” M.Sc. thesis, Dept. of biomedical sc., Univ. of Iceland, Reykjavik, 2016.

[37] Ó. E. Sigurjónsson, “The differentiation Potential of Human Somatic Stem Cells,” Ph.D thesis, faculty of Med., Univ. of Oslo, Oslo, 2006.

[38] M. R. Alison, R. Poulsom, S. Forbes, and N. A. Wright, “An introduction to stem cells,” J. Pathol., vol. 197, no. 4, pp. 419–423, Jul. 2002.

[39] A. I. Caplan, “Mesenchymal stem cells,” Journal of Orthopaedic Research, vol. 9, no. 5, pp. 641–650, Sep. 1991.

[40] L. A. Fortier, “Stem Cells: Classifications, Controversies, and Clinical Applications,” Veterinary Surgery, vol. 34, no. 5, pp. 415–423, Sep. 2005.

[41] “Mesenchymal Stem Cell Research: Some MSCs are More Equal than Others.” [Online]. Available: https://www.promocell.com/cells-in-action/some-human-mesenchymal-stem-cells-are-more-equal- than-others-origins-and-differences/. [Accessed: 11-Jan-2019].

[42] E. Birmingham, G. L. Niebur, P. E. McHugh, G. Shaw, F. P. Barry, and L. M. McNamara, “Osteogenic differentiation of mesenchymal stem cells is regulated by osteocyte and osteoblast cells in a simplified bone niche,” Eur Cell Mater, vol. 23, pp. 13–27, Jan. 2012.

[43] J. Yu et al., “Induced Pluripotent Stem Cell Lines Derived from Human Somatic Cells,” Science, vol. 318, no. 5858, pp. 1917–1920, Dec. 2007.

[44] K. Takahashi and S. Yamanaka, “Induction of pluripotent stem cells from mouse embryonic and adult fibroblast cultures by defined factors,” Cell, vol. 126, no. 4, pp. 663–676, Aug. 2006.

[45] J. W. Haycock, “3D cell culture: a review of current approaches and techniques,” Methods Mol. Biol., vol. 695, pp. 1–15, 2011.

[46] R. Derda et al., “Paper-supported 3D cell culture for tissue-based bioassays,” PNAS, vol. 106, no. 44, pp. 18457–18462, Nov. 2009.

[47] J. Hansmann, F. Groeber, A. Kahlig, C. Kleinhans, and H. Walles, “Bioreactors in tissue engineering- principles, applications and commercial constraints,” Biotechnology Journal, vol. 8, no. 3, pp. 298– 307, Mar. 2013.

[48] F. Zhao, P. Pathi, W. Grayson, Q. Xing, B. R. Locke, and T. Ma, “Effects of oxygen transport on 3-d human mesenchymal stem cell metabolic activity in perfusion and static cultures: experiments and mathematical model,” Biotechnol. Prog., vol. 21, no. 4, pp. 1269–1280, Aug. 2005.

[49] E. Fuchs, T. Tumbar, and G. Guasch, “Socializing with the neighbors: stem cells and their niche,” Cell, vol. 116, no. 6, pp. 769–778, Mar. 2004.

[50] I. Martin, D. Wendt, and M. Heberer, “The role of bioreactors in tissue engineering,” Trends in Biotechnology, vol. 22, no. 2, pp. 80–86, Feb. 2004.

[51] A. Ratcliffe and L. Niklason, “NYAS Publications, Bioreactors and Bioprocessing for tissue engineering,” The New York Academy of Sciences, Jan. 2006.

[52] J. Ge et al., “The size of mesenchymal stem cells is a significant cause of vascular obstructions and stroke,” Stem Cell Rev, vol. 10, no. 2, pp. 295–303, Apr. 2014.

[53] J. F. Alvarez-Barreto and V. I. Sikavitsas, “Improved Mesenchymal Stem Cell Seeding on RGD‐ Modified Poly(L‐lactic acid) Scaffolds using Flow Perfusion,” Macromolecular Bioscience, 2007.

[54] Griffon, Dominique J et al. “A comparative study of seeding techniques and three-dimensional matrices for mesenchymal cell attachment.” Journal of tissue engineering and regenerative medicine 5 3 (2011): 169-79.

[55] S. Kapur and D. Baylink, “Fluid flow shear stress stimulates human osteoblast proliferation and differentiation through multiple interacting and competing signal transduction pathways,” Mar. 2003.

[56] S. C. Cowin, “Bone poroelasticity,” Journal of Biomechanics, vol. 32, no. 3, pp. 217–238.

[57] U. Meyer, H. Wiesmann, and B. Kruse-Lösler, “Strain-related bone remodeling in distraction osteogenesis of the mandible,” Mar. 1999.

[58] L. Meinel et al., “Bone tissue engineering using human mesenchymal stem cells: effects of scaffold material and medium flow,” Ann Biomed Eng, vol. 32, no. 1, pp. 112–122, Jan. 2004.

[59] Synthecon, “Rotary Cell Culture Systems.” [Online]. Available: http://synthecon.com/pages/rotary_cell_culture_systems_13.asp. [Accessed: 02-Dec-2018].

[60] T. J. Goodwin, T. L. Prewett, D. A. Wolf, and G. F. Spaulding, “Reduced shear stress: A major component in the ability of mammalian tissues to form three‐dimensional assemblies in simulated microgravity,” Journal of Cellular Biochemistry, vol. 51, no. 3, pp. 301–311, Mar. 1993.

[61] C. R. Jacobs, C. E. Yellowley, B. R. Davis, Z. Zhou, J. M. Cimbala, and H. J. Donahue, “Differential effect of steady versus oscillating flow on bone cells,” J Biomech, vol. 31, no. 11, pp. 969–976, Nov. 1998.

[62] E. Volkmer et al., “Hypoxia in static and dynamic 3D culture systems for tissue engineering of bone,” Tissue Eng Part A, vol. 14, no. 8, pp. 1331–1340, Aug. 2008.

[63] V. I. Sikavitsas, J. S. Temenoff, and A. G. Mikos, “Biomaterials and bone mechanotransduction,” Biomaterials, vol. 22, no. 19, pp. 2581–2593, Oct. 2001.

[64] A. Ignatius et al., “Tissue engineering of bone: effects of mechanical strain on osteoblastic cells in type I collagen matrices,” Biomaterials, vol. 26, no. 3, pp. 311–318, Jan. 2005.

[65] K. S. Eyres, M. Saleh, and J. A. Kanis, “Effect of pulsed electromagnetic fields on bone formation and bone loss during limb lengthening,” Bone, vol. 18, no. 6, pp. 505–509, Jun. 1996.

[66] R. J. Midura et al., “Pulsed electromagnetic field treatments enhance the healing of fibular osteotomies,” J. Orthop. Res., vol. 23, no. 5, pp. 1035–1046, Sep. 2005.

[67] E. K. Antonsson and R. W. Mann, “The frequency content of gait,” J Biomech, vol. 18, no. 1, pp. 39– 47, 1985.

[68] C. T. Rubin, H. J. Donahue, J. E. Rubin, and K. J. McLeod, “Optimization of electric field parameters for the control of bone remodeling: exploitation of an indigenous mechanism for the prevention of osteopenia,” J. Bone Miner. Res., vol. 8 Suppl 2, pp. S573-581, Dec. 1993.

[69] L.-Y. Sun, D.-K. Hsieh, P.-C. Lin, H.-T. Chiu, and T.-W. Chiou, “Pulsed electromagnetic fields accelerate proliferation and osteogenic gene expression in human bone marrow mesenchymal stem cells during osteogenic differentiation,” Bioelectromagnetics, vol. 31, no. 3, pp. 209–219, Apr. 2010.

[70] J. Halpern et al., “The application of a murine bone bioreactor as a model of tumor: bone interaction,” Clin. Exp. Metastasis, vol. 23, no. 7–8, pp. 345–356, 2006. 56

[71] G. E. Holt, J. L. Halpern, C. C. Lynch, C. J. Devin, and H. S. Schwartz, “Imaging analysis of the in vivo bioreactor: a preliminary study,” Clin. Orthop. Relat. Res., vol. 466, no. 8, pp. 1890–1896, Aug. 2008.

[72] P. H. Warnke et al., “Man as living bioreactor: fate of an exogenously prepared customized tissue- engineered mandible,” Biomaterials, vol. 27, no. 17, pp. 3163–3167, Jun. 2006.

[73] G. H. van Lenthe, H. Hagenmüller, M. Bohner, S. J. Hollister, L. Meinel, and R. Müller, “Nondestructive micro-computed tomography for biological imaging and quantification of scaffold– bone interaction in vivo,” Biomaterials, vol. 28, no. 15, pp. 2479–2490, May 2007.

[74] J. D. Boerckel, D. E. Mason, A. M. McDermott, and E. Alsberg, “Microcomputed tomography: approaches and applications in bioengineering,” Stem Cell Res Ther, vol. 5, no. 6, Dec. 2014.

[75] K. Balto, R. Müller, D. C. Carrington, J. Dobeck, and P. Stashenko, “Quantification of periapical bone destruction in mice by micro-computed tomography,” J. Dent. Res., vol. 79, no. 1, pp. 35–40, Jan. 2000.

[76] R. D. Kapadia et al., “Applications of micro-CT and MR microscopy to study pre-clinical models of osteoporosis and osteoarthritis,” Technol Health Care, vol. 6, no. 5–6, pp. 361–372, Dec. 1998.

[77] H. Grai chen et al., “A non-destructive technique for 3-D microstructural phenotypic characterisation of bones in genetically altered mice: preliminary data in growth hormone transgenic animals and normal controls,” Anat. Embryol., vol. 199, no. 3, pp. 239–248, Mar. 1999.

[78] J. Zeltinger, J. K. Sherwood, D. A. Graham, R. Müeller, and L. G. Griffith, “Effect of pore size and void fraction on cellular adhesion, proliferation, and matrix deposition,” Tissue Eng., vol. 7, no. 5, pp. 557–572, Oct. 2001.

[79] A. S. P. Lin, T. H. Barrows, S. H. Cartmell, and R. E. Guldberg, “Microarchitectural and mechanical characterization of oriented porous polymer scaffolds,” Biomaterials, vol. 24, no. 3, pp. 481–489, Feb. 2003.

[80] “DICOM introduction and overview.” [Online]. Available: http://dicom.nema.org/medical/dicom/current/output/chtml/part01/chapter_1.html#sect_1.1. [Accessed: 19-Nov-2018].

[81] D. L. Logan, A first course in the finite element method, 5th ed. Stamford, CT: Cengage Learning, 2012.

[82] C. Jungreuthmayer, M. J. Jaasma, A. A. Al-Munajjed, J. Zanghellini, D. J. Kelly, and F. J. O’Brien, “Deformation simulation of cells seeded on a collagen-GAG scaffold in a flow perfusion bioreactor using a sequential 3D CFD-elastostatics model,” Med Eng Phys, vol. 31, no. 4, pp. 420–427, May 2009.

[83] G. Kaklamani, D. Cheneler, L. M. Grover, M. J. Adams, and J. Bowen, “Mechanical properties of alginate hydrogels manufactured using external gelation,” J Mech Behav Biomed Mater, vol. 36, pp. 135–142, Aug. 2014.

[84] A. V. Salsac, L. Zhang, and J. M. Gherbezza, “Measurement of mechanical properties of alginate beads using ultrasound,” p. 6, 2009.

[85] Y. Zhu, Z. Dong, U. C. Wejinya, S. Jin, and K. Ye, “Determination of mechanical properties of soft tissue scaffolds by atomic force microscopy nanoindentation,” J Biomech, vol. 44, no. 13, pp. 2356– 2361, Sep. 2011.

[86] E. L. Abreu, M. P. Palmer, and M. M. Murray, “Collagen density significantly affects the functional properties of an engineered provisional scaffold,” J Biomed Mater Res A, vol. 93, no. 1, pp. 150–157, Apr. 2010.

[87] A. Gaspar, L. Moldovan, D. Constantin, A. Stanciuc, P. Sarbu Boeti, and I. Efrimescu, “Collagen– based scaffolds for skin tissue engineering,” J Med Life, vol. 4, no. 2, pp. 172–177, May 2011.

[88] S. H. Cartmell, B. D. Porter, A. J. García, and R. E. Guldberg, “Effects of medium perfusion rate on cell-seeded three-dimensional bone constructs in vitro,” Tissue Eng., vol. 9, no. 6, pp. 1197–1203, Dec. 2003.

[89] B. Porter, R. Zauel, H. Stockman, R. Guldberg, and D. Fyhrie, “3-D computational modeling of media flow through scaffolds in a perfusion bioreactor,” J Biomech, vol. 38, no. 3, pp. 543–549, Mar. 2005.

[90] B. D. Porter, S. H. Cartmell, and R. E. Guldberg, “Effect of media perfusion rate on cell seeded 3D bone constructs in vitro,” in Proceedings of the Second Joint 24th Annual Conference and the Annual Fall Meeting of the Biomedical Engineering Society] [Engineering in Medicine and Biology, 2002, vol. 1, pp. 888–889 vol.1.

[91] “Reynolds Number.” [Online]. Available: https://www.engineeringtoolbox.com/reynolds-number- d_237.html. [Accessed: 30-Oct-2018].

[92] Z. Zhang, L. Yuan, P. D. Lee, E. Jones, and J. R. Jones, “Modeling of time dependent localized flow shear stress and its impact on cellular growth within additive manufactured titanium implants,” J Biomed Mater Res B Appl Biomater, vol. 102, no. 8, pp. 1689–1699, Nov. 2014.

[93] “Stress, Strain and Hooke’s Law - Lesson,” www.teachengineering.org. [Online]. Available: https://www.teachengineering.org/lessons/view/van_cancer_lesson2. [Accessed: 13-Nov-2018].

[94] J. Lovecchio, “Development of an innovative bioreactor system for human bone tissue engineering,” Ph.D thesis, Dep. in Cesena, Univ. of Bologna, Italy, and Reykjavik Univ., Iceland. 2018.

[95] P. M. Hinderliter et al., “ISDD: A computational model of particle sedimentation, diffusion and target cell dosimetry for in vitro toxicity studies,” Particle and Fibre Toxicology, vol. 7, no. 1, p. 36, Nov. 2010.

[96] “Water - Dynamic and Kinematic Viscosity.” [Online]. Available: https://www.engineeringtoolbox.com/water-dynamic-kinematic-viscosity-d_596.html. [Accessed: 20- Dec-2018].

[97] J. Kovácik, “Correlation between Young’s Modulus and Porosity in Porous Materials,” ResearchGate, Jul. 1999.

[98] M. Arnold, A. R. Boccaccini, and G. Ondracek, “Prediction of the Poisson’s ratio of porous materials,” Journal of Materials Science, vol. 31, no. 6, pp. 1643–1646, 1996.

[99] M. Jagodzinski et al., “Influence of perfusion and cyclic compression on proliferation and differentiation of bone marrow stromal cells in 3-dimensional culture,” J Biomech, vol. 41, no. 9, pp. 1885–1891, 2008.

[100] “Wireless pH Sensor - PS-3204: PASCO.” [Online]. Available: https://www.pasco.com/prodCatalog/PS/PS-3204_wireless-ph-sensor/index.cfm. [Accessed: 04-Jan- 2019].

[101] B. D. Porter, A. S. P. Lin, A. Peister, D. Hutmacher, and R. E. Guldberg, “Noninvasive image analysis of 3D construct mineralization in a perfusion bioreactor,” Biomaterials, vol. 28, no. 15, pp. 2525– 2533, May 2007.

[102] “Products | Bellco Glass.” [Online]. Available: http://bellcoglass.com/products/spinner-flasks?page=1. [Accessed: 08-Jan-2019].

[103] “Corning® Disposable Spinner Flasks | Disposable Spinner Flasks | Bioprocess and Scale up | Life Sciences United States Business Site | Corning.” [Online]. Available: https://ecatalog.corning.com/lifesciences/b2b/US/en/Bioprocess-and-Scale-up/Disposable-Spinner-Flasks/Corning%C2%AE- Disposable-Spinner-Flasks/p/corningDisposableSpinnerFlasks. [Accessed: 08-Jan-2019].

[104] W. W. Minuth and L. Denk, “Bridging the gap between traditional cell cultures and bioreactors applied in regenerative medicine: practical experiences with the MINUSHEET perfusion culture system,” Cytotechnology, vol. 68, no. 2, pp. 179–196, Mar. 2016.

[105] “BISS TGT Perfusion Bioreactor Systems.” [Online]. Available: http://www.tissuegrowth.com/prod_bone.cfm. [Accessed: 08-Jan-2019].

[106] “Zellwerk GmbH - Red Point.” [Online]. Available: http://www.zellwerk.biz/prod_biostat.htm. [Accessed: 08-Jan-2019].

[107] “Flexcell International Corporation.” [Online]. Available: http://www.flexcellint.com/. [Accessed: 08- Jan-2019].

[108] “FX-5000 Compression System.” [Online]. Available: http://www.flexcellint.com/FX5000C.htm. [Accessed: 08-Jan-2019].

Appendix A – Bioreactor manual

Bioreactor manual

User manual, version 1.1

6Contents

1. General Information...... 61 Technical information ...... 61 Environmental and electrical specifications ...... 62 System assembly ...... 62 Sterilization of system ...... 62 Connections to Supply ...... 62 Safety Warnings ...... 62 Manual Operating Controls ...... 63 General Maintenance ...... 63

2. Components overview ...... 64 Chamber components: ...... 64 Other components ...... 65

3. Sterilization and assembly ...... 69 Disassemble ...... 69 Assembly for cell culture on scaffolds ...... 70

4. Software overview ...... 74 Initialization ...... 74 Controls ...... 74

General Information

This bioreactor system is an instrument specifically designed for 3D stem cell cultures, growing inside scaffolds. The device is capable of deformation of the scaffold by compression and it provides constant indirect perfusion of media. The compression cycle, perfusion speed and operating temperature can be controlled by the user. The bioreactor chamber includes three main parts and one of them can be custom made to fit the size of scaffolds used. All the chamber parts are made by 3D printing with the MED610 biocompatible hard plastic material. A graphical user interface (GUI) allows user to control parameters of the device. The device is made to be “hands-free” closed system to minimize infection risk. Media can be applied to system via ports on chamber and removed in the same way.

Technical information

Chamber dimensions (w,d,h)* 102x106x90 mm Bioreactor dimensions (w,d,h) 40x40x40 cm Weight of chamber 500 g Applied stimulations Perfusion, compression, temperature Compression axes Uniaxial Configuration 3D scaffolds Number of scaffolds 1-10 (internal plate dependent) Culture area Scaffold size dependent Strain 1-50% uniaxial (scaffold dependent) Optimal perfusion velocity 1-5 ml/min Operating frequency See table below Recommended elastic modulus of scaffold 103 –107 [Pa] Media volume 40-60 mL (must be above perfusion ports)

* width, depth, height

Compression unit: Amplitude of wave (µm) Operating frequency 100 0.5-4 Hz 200 0.5-3 Hz 300 0.5-2 Hz

Environmental and electrical specifications

Electrical Input 12V DC Current Rating 2A Environmental Maximum operating temperature - 40°C Conditions 0% - 100% Relative Humidity Data connections USB to custom made computer for the device Control system Custom made Arduino circuit board Graphical User Interface (GUI) Custom made – Easy to use

System assembly

An overview of system components can be found in chapter 2 in this manual. The Bioreactor consists of main support unit, the chamber and connecting tubes between them. Other items such as gas tank and sterilizing hood are also necessary. First step is to measure the dimensions of scaffolds used and make sure that the internal plate has enough space for them to fit. Follow the sterilization protocol for full assembly of system.

Sterilization of system

See chapter: Sterilization

Connections to Supply

USB-A connector connects from main support unit to computer. Connectors from main support unit and computer must be connected to power outlet. Disconnect the USB connector to computer before starting it up. Then connect it when system has booted.

Safety Warnings

This equipment must be used in accordance with the procedures outlined in this manual to operate correctly.

Manual Operating Controls

There are several parameters which user can change via GUI, they are: • Perfusion flow rate • Compression cycle rate and amplitude • Input and output media flow speed • Temperature control

The PH value is controlled manually by hand by adjusting the flow of gas: - If PH > 7.4 → increase bubble count - If PH < 7.4 → decrease bubble count - 7.2-7.6 is optimal value - Should be around 5-6 bubbles/minute

Controls are explained in detail in the software overview chapter.

General Maintenance

The equipment does not require maintenance other than cleaning and sterilizing. Please see components overview in next chapter for cleaning and sterilization details.

Components overview

Chamber components:

Chamber Description: Housing for cell culture, provides closed sterile environment

Material: MED610, agilus30 rubber

Cleaning: Bacteria killing soap, ethanol, UV

Internal plate Description: Designed to hold scaffolds in

place between its “teeth”.

Material: MED610

Cleaning: Bacteria killing soap, ethanol, UV

Molding/support plate Description: 2 purposes. Its main purpose

is to hold the internal plate in place when inside chamber. It can also be used as support when molding hydrogel scaffolds.

Material: MED610

Cleaning: Bacteria killing soap, ethanol, UV Compression plate Description: Applies compression to scaffolds, powered by motor

Material: MED610

Cleaning: Bacteria killing soap, ethanol, UV

Lid Description: Lid on box. Keeps chamber environment isolated

Materials: MED610, agilus30 rubber

Cleaning: Bacteria killing soap, ethanol, UV

Motor Description: Applies compression

Cleaning: Only clean screw part with bacteria killing soap, ethanol and UV

Motor rubber Description: Damping unit between motor and chamber.

Cleaning: Bacteria killing soap, ethanol, UV

Other components

Main support unit (Metal case) Description: Holds components in place.

Peristaltic pumps (3x) Description: Produces peristaltic flow for perfusion and input/output media

Cleaning: Only essential to clean tube inside housing with bacteria killing soap, ethanol and UV

PH sensor Description: Reads PH value of media

Cleaning: Bacteria killing soap, ethanol, UV

Temperature sensor Description: Measures temperature of media

Cleaning: Bacteria killing soap, ethanol, UV

Control unit Description: Main circuit board of bioreactor

Cleaning: don’t clean!

Open pipes Description: Connectors for media or gas transport. Female to male.

Cleaning: Bacteria killing soap, ethanol, UV

Microclave pipes Description: Connectors for media or gas transport. Male to male Microclave connectors.

Cleaning: If new from bag, no need for cleaning, otherwise use bacteria killing soap, ethanol, UV

Filters Description: Keeps environment sterile but allows free flow of gasses. Use new every run.

White connectors - Microclave Description: Connects to Spyros connector on narrow end and M/F pipe on other. It remains closed when not attached so it keeps environment sterile

Cleaning: Bacteria killing soap, ethanol, UV Blue connectors - Spiros Description: Connects to Microclave connector on thick end and screw mechanism on box on other. It remains closed when not attached so it keeps environment sterile.

Cleaning: Bacteria killing soap, ethanol, UV Media bag Description: Stores input or output media and can be inserted into screw mechanism on chamber. At Blood bank they are made from fusing together pipe with Microclave head and media bag pipe.

Gas tank Description: Provides correct gas combination for osteogenic cell cultures; 5% CO2, 20% O2 and 75% N2.

Bubble counter Description: Visual representation of flow of gas from gas tank. Should be 5-10 bubbles/minute when operating.

Cleaning: Bacteria killing soap, ethanol, UV

Heating element Description: Sits under chamber and heats up the media inside it.

Sterilization and assembly

Sterilization of Bioreactor. Utilities needed are cell culture hood, antibacterial soap and UV light. Items needed: Sterilized gloves

Disassemble

- Unscrew two screws to release motor with screwdriver o Unscrew big motor screw manually to release it from chamber o Take out all components - Unplug PH sensor - Unplug temperature sensor (back cover has to be removed and connector should be disconnected straight from control unit circuit board). - Remove rubber square from motor and clean it o Wash big screw on motor with soap (only the screw!) - Use bacterial killing soap to clean chamber components and the two small screws - Clean bubble counter o Good to clean with paper or small brush on the inside - Clean sensors (PH and Temp) - Clean tubes of rotators (two short and one longer) - Clean all tubes needed.

- Insert all components into hood (spray with ethanol first well, inside and outside (not inside of motor)). See images above - Turn on the UV light inside the hood and leave it on for 20 minutes.

Assembly for cell culture on scaffolds

Bring scaffolds, and tools needed into hood and follow these steps inside sterile the hood. Chamber preparation: - Insert the internal plate and supporting plate into chamber - Place motor into box through the back hole o Screw metal wheel in place on compression plate - Screw blue connectors (6x) on every screw mechanism of chamber (do not screw the white connectors on). This seals the open hole. - Place internal plate, and support plate in chamber. - Insert motor screw through hole on back of chamber (make sure that the rubber square is in between. The compression plate must be inserted simultaneously. - Secure the motor in place with two screws to back of chamber. - Now put temperature and PH sensors in the chamber and put the lid on. - Now every entry of air should be enclosed. Set the chamber aside but keep it in the hood.

Metal case preparation: - Start on motor parts. Find three short matte pipes (not transparent), two are smaller, one is bigger. - Screw blue connectors on shorter pipe ends (4x connectors) - Screw white connectors on longer pipe ends (2x connectors) - Take out of hood and insert pipes into pump housing units. - Back in hood: Put right connectors on open pipes to close them (to keep them sterile while not inside hood). - Here is a list of connecting pipes needed: o Two double open pipes (two connected together for perfusion) o One with white connector on one end and filter on other end (for ventilation) o Two closed pipes (connecting media bags to chamber) o Two matte pipes for gas. - Make sure before taking pipes out of hood that they are closed to keep them sterilized. - Assemble all pipes on metal case.

Scaffold preparation: (items recommended: 6 well plate, 24 well plate, pipette (20-1000 µL (yellow one)), pipette heads, cell media). Make sure you have cells ready to be seeded by following the seeding protocol made by the Blood bank. Cells should be in a 15mL tube and suspended in 1mL media if seeding protocol was followed. - Put on new pair of sterilized gloves - If using wet scaffolds: Put small amount of fresh media in one well of 6 wells plate, use that to dip the wet scaffolds with sterilized tools before moving it to 24 well plate. - Use sterilized tools to move scaffolds to any wells in 24 well plate, mark them properly. - Divide the cells between the scaffolds (or if you want specific count on every scaffold you should count the cells with counting protocol and divide accordingly). - Use the pipette to slowly feed the cells to the scaffold (ca. 30 sec. per scaffold) and then incubate for 5 minutes. This gets cells into pores of scaffold - Do step above for approx. 1 hour (ca. 5-6x times). - Open the chamber and place scaffolds in between the “teeth” of internal plate with sterilized tools (or with hands, wearing sterilized gloves). 72

- Turn the motor screw manually (to move compression plate closer) until scaffolds are approximately 1 mm from compression plate. - Put around 50-60 mL media into chamber (so it covers the perfusion holes). - Place temperature and PH sensors in place and then put the lid on. - Now the chamber is ready to be taken out of the hood.

Hydrogel molding

The internal plate design offers a solution for hydrogel scaffold molding. Fix the support plate on top of internal plate (see images below), to be able to mold hydrogel scaffolds in between the teeth of the internal plate while using the support plate as a cover.

System Initialization - Connect temperature sensor back to control unit (and close the back cover of metal case). - Connect PH sensor - Plug all connectors on to chamber according to image below (conventional setup)

- Turn on the computer, Password is: Biomedical2017 (ignore annoying error messages, press cancel) - Open software in: start/PaoloGargiulo/Documents/Bioreactorapp/lastversion - For software instructions, see next chapter.

Software overview

A specially designed graphical user interface was developed for the bioreactor system. Many parameters can be defined by user. The interface allows to control easily the actuators through a series of virtual on/off switches and adjustable bars, simultaneously allowing the constant monitoring of each sensor.

Initialization

Make sure the USB connector from metal base back is connected to computer and open the bioreactor software located in /documents/Bioreactorapp/Lastversion. First comes up a settings screen where following changes must be made: - Set input to: COM6 - Change sample frequency to 1Hz - Choose destination file for output data (best to create empty text file beforehand)

Other setting can be left as default. Next open the controls tab and press arrow in top left corner so it becomes filled. Next turn on “Send commands” button and then the “read data” button. They should light up green.

Controls

Compression stimulus can be applied by setting the displacement [μm] and the frequency [Hz] of the compression plate integrated within the compression unit; two buttons (Up, Down) can be used to align the compression plate, rather, to allow the contact between the cylindrical scaffold units and the pistons of the compression unit. Perfusion can be applied and tuned changing the flow intensity [ml/min] related to the dedicated peristaltic pump. Automatic media replacement can be performed setting the “inflow” and “outflow" parameters related to the dedicated peristaltic pumps. The desired culture temperature can be set through the “Set Temp” graphic element. Temperature [°C], Pressure [kPa] and pH [-] can be monitored by dedicated scalebar. Two arrays (“Serial In”, “Serial Out”) are useful to check constantly if the right input/output are correctly sent/received to/from the CU. A “STOP” button is present to allow the interruption of all operations.

Image below shows the graphical user interface.

Appendix B – Cleaning of chamber for bioreactor system (at RU)

The following instructions apply for cleaning and getting the bioreactor box ready to use for cell therapy. Make sure your workspace is clean and free of residue from other materials. Waterjet cleaning First step is to remove as much supporting material as possible with the waterjet machine in the electronic room. How to: 1. Turn the knobs on the sink as shown in the picture to make sure there is flow to the waterjet machine. Upper one should be closed (turned perpendicular to stream direction) and bottom knob should be parallel to stream direction).

2. Use hot water because it makes it easier to remove support material. 3. Turn on waterjet with switch, enter waterjet by putting on the rubber gloves and use the footswitch to start applying high pressure water stream.

4. Clean printed parts thoroughly with 10 rounds of rinses on every surface of chamber. Take extra care around the thin surfaces (around screw mechanism) but it should nonetheless be able to withstand the waterjet pressure even from close range. 5. Remove parts from waterjet. Rinse in the sink with water. Next step is further cleaning with salt solution.

Salt solution The salt solution removes the remaining layers of support material. Always wear protective gloves when working with salt solution and be ready to clean up immediately if a drop spills on table because the material will leave mark. 1. Salt ratio: a. 2% Sodium Hydroxide (SH) b. 1% Sodium Metasilicate (SM) 2. Obtain weight scale that has accuracy of at least 1 gram. 3. Examples of mixtures: a. If using 6L water bath, add: i. 120g SH ii. 60g SM b. If using 4L water bath, add: i. 80g SH ii. 40g SM c. If using 2L water bath, add: i. 40g SH ii. 20g SM d. If using 1L water bath, add: 78

i. 20g SH ii. 10g SM 4. Never add water to salt, always the other way around: Add the salt slowly into water bath while stirring. 5. The salt causes rise in temperature of the water, wait until it has reached room temperature. 6. Place chamber and other parts in the solution. 7. Start with chamber cleaning, place it in the solution: a. Items recommended (see image): i. Brush (e.g. toothbrush) ii. Brush with less than 2mm diameter wire iii. Wire 1mm diameter iv. Wire 0.5mm diameter b. Start by threading thin wire through the holes, and gradually work all the way through all holes c. Start using thicker wire and then brush. It can take a while to reach all the way through long holes d. Clean every surface with brush, and clean extra well the inside surface. e. Total time of chamber in solution should not exceed 45 minutes (thin walls might get too eroded). 8. Clean remaining parts with brush and make sure there is no support material in small cavities 9. After being cleaned thoroughly, take box out of solution and wash in sink 10. Remaining support material can be washed off in waterjet if necessary

Light treatment Bright light for few hours can make material more transparent and it removes undesired yellow tint from chamber. 1. Plug in the TUBORG box and make sure the thermometer is inside the box

2. Turn on the light 3. Place the components inside the box and close it 4. Monitor the temperature and have it in the range of 35-45 °C 5. Leave it for 2 hours up to 24 hours depending on how much tint you want to be removed.

Appendix C – Cell experiments in bioreactor and other

Cell experiments in bioreactor

Standard protocol procedures of the Bloodbank of Iceland and Platome Biotechnology were followed to grow, harvest and seed mesenchymal stem cells and to perform osteogenic differentiation. In order to carry out a successful run every time in the bioreactor, protocols were made for both how to sterilize it and how to seed scaffolds with cells and place them in the chamber. These protocols include clear step by step instructions, so others can use the bioreactor properly. These protocols can be found in appendix A (in the bioreactor manual). After these protocols were followed, test runs were made with different setups. The idea was to have four scaffolds in osteogenic differentiation in the bioreactor and simultaneously having two in static culture as reference. However, the bioreactor runs did not work out as desired. Three test runs were made with MSCs infused in bioactive glass scaffolds. Here is a description of them. The original plan was to have data from the bioreactor parallel to the static culture. However, they did not go according to plan due to problems with the bioreactor system. The scaffolds used were bioactive 47.5B glass scaffolds. Their composition is: 47.5% SiO2, 20% CaO, 2.5% P2O5, 10% MgO, 10% Na2O and 10% K2O.

Run 1

The setup for this run can be seen in figure below. This is the original setup of the system. Here the gas was pumped straight into the chamber. This allows the PH regulation to take place in the media where the scaffolds are. 1,5 million cells were divided onto six different scaffolds (250 thousand each), three graded and three monopored scaffolds. Four of them (two graded and two monopores) went into the bioreactor and the other two (one of each) into static culturing process. This run was unsuccessful because the pH value of the media was not stable enough. The cells died because of lack of gas reaching the chamber. Next run carried out tried to resolve that problem.

Run 2

New setup was carried out as seen in figure below An issue with pH control led to the idea of having the gas introduced into a reservoir media located in a bag in the perfusion cycle. This run was unsuccessful because of few reasons. First, when the perfusion was turned on, the reservoir media bag gradually emptied, and it filled the chamber with media, resulting in leakage. Secondly, the power went off the building during night and since the bioreactor is dependent on having constant electrical input, it turned off.

Run 3

Another attempt was carried out with the same configuration as in run 1. This run was also unsuccessful, mainly because of broken pipes. After post inspection of the system, it turned out that connecting pipes were worn out due to frequent usage. This resulted in gas not reaching the chamber. It leaked out before reaching chamber, resulting in increased PH, that eventually results in dead cells.

Static culture run

A 2D static culture was carried out next to the bioreactor runs. This were supposed to be control for the bioreactor runs. Cells were seeded onto two scaffolds, one graded and one monopore, with methods described in appendix A. They were then submerged in osteogenic differentiation media. Media was changed on Mondays, Wednesdays and Fridays for two weeks, a total of 14 days. After that the scaffolds were placed in formaldehyde for cell fixation for five days. Then they were µCT scanned for extra cellular matrix growth evaluation. The scanner malfunctioned and required maintenance, so because it was time consuming, unfortunately there was no time to analyze the data and obtain results.

Attempts making chamber more transparent

A few experiments were made on the 3D printed chamber and its components, to make it become more transparent. Main reason for this is to make it easier to observe and visualize what is going on inside the chamber while running experiments. During printing, the support material (waste material from 3D printing) attached to chamber made the walls of it matte and less transparent. Attempts using dry sanding, wet sanding and epoxy resin were carried out. The main method used was dry sanding with rough sandpaper, followed by wet sanding with gradually finer sandpaper using circular motion. After many tests with different types of sandpapers, the results were worse than initially. Epoxy-resin solution was applied on surface with the purpose of filling in the micro holes on the surface that caused matte finish. This fixed the sanding errors but eventually the surface became the same as when started. The conclusion is that the 3D printed MED610 material could not become clear due to the micro structure of the material. 84

µCT scans of scaffolds

The six bioactive glass scaffolds were sent to sterilization and after receiving them, they were dry scanned in a µCT scanner located at Nýsköpunarmiðstöð Íslands (Figure below). The scaffolds were seeded, and osteogenic differentiation was performed for 14 days in static culture. After that, the scaffolds were submerged in water and scanned again. Only half of each scaffold was scanned because of large file size, computer processing power demand and excessive memory usage. On average, around 1000 dicom files were obtained per scaffold. By using the software Mimics by Materialize, a 3D image was constructed. By analyzing the histogram (different gray level of different materials in scans), the extra cellular matrix secreted by cells during osteogenic differentiation can be quantified. But due malfunction of the scanner there was not enough time to complete this part.

Appendix D – Matlab Code

Young’s modulus, Poisson’s ratio and density of scaffold as a function of porosity: clc clear all

% Bulk values E0 = 72.9E9; v0 = 0.2679; p0 = 2640;

% Percentage porosity, 0 -> bulk, 0,5 -> 50% porous, 1 -> empty space x = linspace(0,1,101);

% Young's modulus formula E = E0*(1-x).^2/10E8;

% Poisson's ratio formula s = 1./(1+exp(-100.*(x-0.4))); v1 = 0.5-(1-x.^(2/3)).^(1.21)./(4.*(((1-s).*((3-5.*x).*(1-x))./(2.*(3- 5.*x).*(1-2.*v0)+3.*x.*(1+v0)))+s.*((1-x)./(3.*(1-v0)))));

% Density formula p = -p0*x+p0;

% Plotting figure subplot(1,3,1) plot(x,E,'linewidth',2) title("Young's modulus as a function of porosity") xlabel('porosity, p [%/100]') ylabel("Young's modulus, E [GPa]") grid on subplot(1,3,2) plot(x,v1,'linewidth',2) title("Poisson's ratio as a function of porosity") xlabel('porosity, p [%/100]') ylabel("Poisson's ratio, v [ ]") grid on subplot(1,3,3) plot(x,p,'linewidth',2) title("Density as a function of porosity") xlabel('porosity, p [%/100]') ylabel('Density, d [kg/m^3]') grid on