ABSTRACT DESIGN, FABRICATION, and ANALYSIS of POLYMER SCAFFOLDS for USE in BONE TISSUE ENGINEERING by Joshua Minton Bone Tissue

ABSTRACT

DESIGN, FABRICATION, AND ANALYSIS OF POLYMER SCAFFOLDS FOR USE IN BONE TISSUE ENGINEERING

by Joshua Minton

Bone tissue engineering is an emerging field that seeks to improve the treatment of bone defects by restoring the functions of bone using the body’s natural healing processes. Polymer scaffolds seeded with osteoblast and growth factors is one technique that has shown the potential to speed the healing process and decrease the rehabilitation time from bone defects. The goal of this study is to create viable polymer/ceramic scaffolds through melt processing of polycaprolactone and hydroxyapatite and using polyethylene oxide as porogen. The results of this study show that melt processing of these materials is an effective method for creating stable scaffolds. The properties of these scaffolds can be altered by changing several factors including polymer ratio, ceramic and salt content, and the pressure applied during the fabrication process. Biological analysis shows that the scaffolds seeded with MC3T3-E1 cells are capable of facilitating cell attachment and proliferation in vitro over time.

DESIGN, FABRICATION, AND ANALYSIS OF POLYMER SCAFFOLDS FOR USE IN BONE TISSUE ENGINEERING

A Thesis

Submitted to the

Faculty of Miami University

in partial fulfillment of

the requirements for the degree of

Master of Science

Department of Chemical, Paper, and Biomedical Engineering

Joshua Minton

Miami University

Oxford, Ohio

2013

Advisor ______

Dr. Azizeh Yousefi Moshirabad

Advisor ______

Dr. Paul James

Reader ______

Dr. Paul Urayama

Reader ______

Dr. Jason Berberich

Table of Contents

Title Page i Table of Contents ii List of Tables iv List of Figures v Acknowledgements vi

Chapter 1: Introduction 1 1.1 Bone Structure and Function 2 1.2 Traditional Repair Methods 3 1.3 Tissue Engineering 4 1.3.1 Solvent Casting/Particulate Leaching 8 1.3.2 Gas Foaming/Particulate Leaching 8 1.3.3 Melt Processing 8 1.4 Materials 10 1.4.1 Polymers 10 1.4.2 Ceramics 12 1.5 Scaffold Fabrication 13 1.6 Porosity and Interconnectivity 14 1.7 Mechanical Properties 15 1.8 Cell Attachment and Growth 15 1.7.1 Scanning Electron Microscopy 17 1.7.2 Proliferation 18 1.9 Project Statement 19 1.8.1 Hypothesis 19 1.8.2 Objectives 20 1.10 Thesis Outline 20 1.11 References 22

Chapter 2: Design and Fabrication of Polymer/Ceramic Scaffolds for Bone Tissue 27 Engineering 2.1 Abstract 27 2.2 Introduction 27 2.3 Scaffold Materials 29 2.4 Materials 31 2.5 Scaffold Characterization 32 2.6 Statistical Analysis 32 2.7 Results and Discussion 32 2.7.1. Effect of Polymer Ratio 32 2.7.2. Effect of HA concentration 36 2.7.3. Effect of Pressure 38 2.8 Conclusions 39

2.9 References 41

Chapter 3: Polymer/Ceramic Scaffolds for Bone Tissue Engineering: Fabrication, 43 Analysis, and Cell Growth 3.1 Abstract 43 3.2 Introduction 44 3.3 Materials and Methods 46 3.3.1. Materials 46 3.3.2. Fabrication of PCL and PCL/HA scaffolds 47 3.3.3. Characterization of porous scaffolds 47 3.3.4. Cell Culture 48 3.3.5. Proliferation Assay 48 3.3.6. Preparation for SEM imaging 48 3.3.7. Statistical Analysis 49 3.4 Results and Discussion 49 3.4.1. Characterization of porous scaffolds 49 3.4.2. Proliferation 54 3.4.3. SEM 54 3.5 Conclusions 55 3.6 References 56

Chapter 4: Significance and Future Work 59 4.1 Significance 59 4.2 Future Work 60

iii

List of Tables Table 1.1: Scaffold fabrication techniques in tissue engineering applications 5 Table 1.2: Methods of preparation of porous scaffolds and their characteristics 7 Table 1.3: Properties of commonly-used biomaterials 11 Table 1.4: Properties of PCL and PEO 12 Table 1.5: Optimum pore size for bone regeneration 14 Table 1.6: Optimum parameters for grinding 1 gram of polymer 15 Table 1.7: Comparison of different SEM sample preparation methods 17 Table 2.1: Properties of PCL and PEO 30 Table 2.2: Mechanical properties of human tissues 31 Table 2.3: T-test results of the various salt concentrations and particle sizes tested 35 Table 2.4: T-test results of the mechanical properties of the scaffolds with different HA 38 Table 2.4: concentrations Table 3.1: Mechanical properties of human tissue 45 Table 3.2: Properties of PCL and PEO 46 Table 3.3: Results of physical characterization of scaffolds 52 Table 3.4: Porosity information determined by microCT analysis 53 Table 3.5: Pore size distribution determined by microCT analysis 53

List of Figures

Figure 1.1: Stages of bone repair 3 Figure 1.2: Fabrication of porous scaffolds by gas foaming/particulate leaching 8 Figure 1.3: SEM micrographs of NaCl particles (a) before mixing and (b) after mixing 9 Figure 1.4: SEM images of the surface of PCL/PEO 50/50 (% vol) sample 9 Figure 1.5: The effect of wt.% HA on the strength of HA/polymer composites 13 Figure 2.1: Average Young’s modulus for scaffolds with various PEO/PCL polymer ratios 33 Figure 2.2: Average calculated porosity for scaffolds with various polymer ratios 33 Figure 2.3: Average Young’s modulus for the scaffolds with various salt particles sizes 34 Figure 2.3: and concentrations Figure 2.4: Average calculated porosity for the scaffolds with various salt particles sizes 35 Figure 2.4: and concentrations Figure 2.5: SEM image of a PEO/PCL 60/40 scaffold with 20% salt with particle sizes less 36 Figure 2.5: than 100 µm Figure 2.6: SEM image of PEO/PCL scaffold with 40% salt with particle sizes between 36 Figure 2.6: 150 µm and 250 µm Figure 2.7: Average Young’s modulus for the scaffolds with various HA concentrations 37 Figure 2.8: Average calculated porosity for the scaffolds with various HA concentrations 37 Figure 2.9: SEM image of PEO/PCL 60/40 scaffold 38 Figure 2.10: SEM image of PEO/PCL 60/40 scaffold with 20% HA 38 Figure 2.11: Average Young’s modulus for the scaffolds with various HA concentrations 39 Figure 2.11: and pressure level Figure 2.12: Average calculated porosity for the scaffolds with various HA concentrations 39 Figure 2.12: and pressure level Figure 3.1: SEM image of PCL scaffold 50 Figure 3.2: SEM image of PCL/HA scaffold 50 Figure 3.3: Elemental mapping images of PCL/HA scaffolds 50 Figure 3.4: Overlaid elemental mapping images of PCL/HA scaffolds 50 Figure 3.5: Thermogravimetric analysis of PCL/HA scaffolds fabricated with 20% HA 51 Figure 3.6: Stress-strain curve of PCL/HA scaffolds 52 Figure 3.7: Pore size distribution of PCL and PCL/HA scaffolds 53 Figure 3.8: DNA content of PCL scaffolds at various time points 54 Figure 3.9: SEM images of PCL/HA scaffolds 55 Figure 3.10: SEM images of PCL scaffolds 55

Acknowledgements

There are several people I would like to acknowledge and thank for their help and guidance throughout my time at Miami and my through my Master’s research. I would not be where I am without their support and guidance – my advisors, committee members, faculty and staff at Miami University, and fellow graduate and undergraduate students.

It has been my privilege to work with Dr. Yousefi, my advisor, for the last three years in both undergraduate and graduate research. Dr. Yousefi has been so helpful and kind throughout the whole time and I have enjoyed the experience immensely. I have no doubt that the knowledge and skills I have learned under her supervision will continue to serve me well in my career and future endeavors.

I would also like to thank my co-advisor, Dr. James, who has been very patient and helpful guiding me through the biological aspects of my research that were new and challenging to me and for sharing his lab and resources so freely.

I would like to thank Dr. Urayama and Dr. Berberich for agreeing to be on my committee and for being so accommodating with their schedules. I would also like to thank Dr. Saul for being on my committee for my proposal and for his insight and assistance during my proposal and with my final work.

I would like to thank Doug Hart for being a constant source of support and encouragement throughout my time at Miami, for always going above and beyond to help in any way he could, and for being one of the best friends I could have, especially on the bad days.

I would like to thank the Instrumentation lab and especially Barry Landrum and Jayson Alexander for their inventive ideas and hard work that designed and built many of the tools I used in my research. Much of my work would not have been possible without their help.

I would like to thank Matt Duley and Dr. Edelmann from the Electron Microscopy facility for teaching me how to prepare and image my samples with SEM and for helping me find solutions to the obstacles I faced along the way.

I would like to thank Rosa Akbarzadeh, Cara Janney, and Carlie Focke for their immense help in the lab. So much work has been put into this research and I could never have done it alone. Their help and support has made this all possible.

Chapter 1

1. Introduction

Polymer scaffolds have a variety of applications in today’s medical field. For example, tissue regeneration and cardiac function after myocardial infarction can be improved by an injection of primary skeletal myoblasts. However, these cells have a high death rate once injected, which greatly decreases the effectiveness of the treatment. One solution to this problem is to use polymer scaffolds that have been found to improve cell viability in vivo. [1] Polymer scaffolds have also proven useful in gene therapy, a technique for treating both acquired and hereditary diseases. Gene therapy has shown great promise in treating hemophilia B and other hereditary plasma protein deficiencies; however, the same problem of survival rate in vivo persists. A potential solution to this is the use of polymer scaffolds which have been shown to support long term cell growth and protein delivery in vivo and have been used to cure hemophilic mice for over 12 weeks. [2,3]

The focus of this project was on the use of polymer scaffolds in bone tissue engineering. The goal of bone tissue engineering is to improve traditional repair methods by using biodegradable constructs that hasten the body’s natural repair mechanisms. To restore the biomechanical function of tissue, scaffolds must be designed to match the structural and mechanical properties of the target tissue. [4] To address this challenge, current efforts are focused on designing biomimetic scaffolds. [5, 6] Producing scaffolds with biomechanical properties similar to native bone has many potential advantages upon in vivo implantation. The stiffness would be sufficient to withstand immediate weight bearing. Moreover, it would mean faster rehabilitation for patients.

The effectiveness of tissue-engineering scaffolds is affected by porosity, interconnectivity of the pores, and mechanical properties of the chosen scaffold material. It has been found that porosity and interconnectivity throughout the scaffold are required to promote cell loading and migration, tissue and vasculature growth, and adequate diffusion of interstitial fluid. A major problem with traditional methods of fabricating these scaffolds is that they result in a random pore structure, leading to poor permeability and interconnectivity. [7]

The goal of this project is to develop a method of fabricating polymer scaffolds that will lead to controlled pore sizes and fully interconnected pores while avoiding the use of toxic solvents. The scaffolds will be made through melt processing, which involves heating the polymer powders above their melting temperatures. Poly(ethylene oxide) (PEO) and polycaprolactone (PCL) will be the primary polymers used due to their low melting temperatures (below 70C). The porosity and interconnectivity of the scaffolds was achieved through dissolving PEO in water in the combined PEO/PCL scaffolds. In addition, in bone tissue engineering there is a need for osteoconductive scaffold materials. Incorporation of ceramic particles, e.g. hydroxyapatite (HA), can result in materials with the desired mechanical properties and better biocompatibility and osteoconductivity. To this end, micro-HA was incorporated into the scaffold formulation at target concentrations of 20% and 40% w/w [2]. The produced scaffolds were seeded with osteoblastic cells in cultured medium. Cell attachment and cell growth were evaluated in vitro up to 21 days.

1.1 Bone Structure and Function

Bones are specialized connective tissues that provide both mechanical support and protection for the body. They are comprised of various cell types and an organic matrix. Calcium minerals, mainly calcium and phosphate in the form of hydroxyapatite, provide extra strength to this matrix. Bone is able to undergo some level of regeneration in certain circumstances when there is damage to the tissue. The extent of this regeneration depends greatly on the size of the defect. Non critical size defects heal spontaneously but critical size defects will not be able to completely heal. Nutrition, diffusion, and vascularization are other factors that influence the bone’s healing process. Various sources of cells including mesenchymal stem cells (MSCs) and various signals also play a role in the formation of new bone tissue. [8] Figure 1.1 shows the stages of bone regeneration and healing.

Figure 1.1: Stages of bone repair. a) Blood clot formation; b) Substitution of blood cells by a matrix-rich repair tissue; c) Invasion of blood vessels; d) Organization of predetermined osseous tissue; e) Trabeculae formed by osteoblasts; f) Final repair tissue. [8]

1.2 Traditional Repair Methods

Tissue/organ repair has been the ultimate goal of surgery since ancient times, and has traditionally taken two forms: (i) tissue grafting and organ transplantation, and (ii) alloplastic or synthetic material replacement [9]. When bone tissue is unable to heal itself naturally, a substitutionary material must be used to fill the bone defect. The current gold standard is autogenous bone grafting. This involves filling the defect with bone tissue taken from another place in the body. The complication rate of this treatment can be as high as 30% and may include

donor site morbidity, pain, prolonged hospitalization and rehabilitation, increased risk of deep infection, hematoma, and inflammation. [10-16] Another option is the use of bone tissue from other humans (typically cadavers), which is called allografting. While allografts have been used for years, they have a risk of donor-to-recipient infection (as high as 13%) [17], disease transmission, and host immune responses. [18,19]

1.3 Tissue Engineering

Tissue engineering emerged in the early 1990s to address the limitations of tissue grafting and alloplastic tissue repair. [20] The concept is to transplant a biofactor (cells, genes and/or proteins) within a porous degradable material known as a scaffold. In functional tissue engineering, which aims for regenerating a load-bearing tissue, scaffolds should balance the temporary mechanical function with mass transport to aid biological delivery and tissue regeneration [20]. Fabricating scaffolds with precise porous architecture can help in understanding the role of the scaffold structural parameters on its mechanical properties and mass transport efficiency. The fabrication method used in creating scaffolds has a critical role in shaping the microstructure and morphology of the pores. There are several methods for fabricating porous scaffolds, each with its own disadvantages. Table 1.1 lists the scaffold fabrication techniques used for tissue engineering applications. [21] Table 1.2 lists the advantages and disadvantages of several of these techniques. [9]

Table 1.1: Scaffold fabrication techniques in tissue engineering applications. Adapted from Brahatheeswaran

Dhandayuthapani, Yasuhiko Yoshida, ToruMaekawa, and D. Sakthi Kumar. (2011)

Table 1.2: Methods of preparation of porous scaffolds and their characteristics. Adapted from Hollister, S. J. (2005)

1.3.1 Solvent Casting/Particulate Leaching

Solvent casting/particulate leaching is a technique that involves dispersing mineral (NaCl) or organic particles (saccharose) in a polymer solution. Casting or freeze-drying is then performed as a dispersion process in order to produce porous scaffold. This basic technique has been used to create scaffolds using various polymers but while the approach is relatively simple, there are disadvantages such as the use of toxic solvents and residual solvent that remains in the scaffold. [22-24]

1.3.2 Gas Foaming/Particulate Leaching

Gas foaming is a technique used under high pressure that uses the CO2 saturation of polymer discs, through exposure to high-pressure CO2. Figure 1.2 shows a diagram of this process which combines gas foaming and particulate leaching. The process involves dispersing salt particles in a polymer gel paste, casting on a Teflon mold for solvent evaporation, immersing in water for gas foaming/salt leaching, and finally freeze drying to produce the final scaffold. [25]

Figure 1.2: Fabrication of porous scaffolds by gas foaming/particulate leaching [25]

This technique allows for a high degree of porosity in the final scaffold as well as producing pores with a pore size in the range of 100 µm. [26] However, this technique results in scaffolds with low mechanical strength and poorly defined pore structure. [27]

1.3.3 Melt Processing

Melt processing involves the use of two immiscible polymers mixed together, which fuse to form an interconnected structure when exposed to temperatures above the melting point of the two

polymers. [28,29] Generally, this co-continuous structure occurs when there is a 40:60-60:40 (vol%) composition of the mixture. [30] In this state, both polymers are continuous in the structure. This allows for the formation of an interconnected pore network if one of the polymers can be leached from the scaffold using a solvent.

Classic melt processing involves the use of NaCl and/or a polymer as a porogen as well as the use of an extruder. While the use of an extruder does provide a very well-blended, homogeneous mixture, it also can lead to salt particle breakdown during mixing (as shown in the scanning electron microscopy (SEM) images in Figure 1.3a and 1.3b), which results in small and inconsistent pore sizes. In addition, the channels generated by the interpenetrated polymer network are usually too small for most tissue engineering applications (Figure 1.4a and 1.4b). The process presented in this project attempts to address several of the most common challenges and disadvantages of fabricating these scaffolds.

Figure 1.3: SEM micrographs of NaCl particles Figure 1.4: SEM images at two different (a) before mixing and (b) after mixing magnifications of the surface of PCL/PEO 50/50 (%vol) sample after dissolution of the PEO phase

1.4 Materials

In biomedical applications, there are several criteria for selecting the materials as biomaterials, including their material chemistry, molecular weight, solubility, shape and structure, hydrophilicity/hydrophobicity, water absorption, and degradation/erosion mechanism.

1.4.1 Polymers Polymers have been widely used as biomaterials for the fabrication of medical device and tissue- engineering scaffolds. In biomedical applications, there are several criteria for selecting the materials as biomaterials, including their material chemistry, molecular weight, solubility, shape and structure, hydrophilicity/hydrophobicity, water absorption, and degradation/erosion mechanism. Table 1.3 lists the properties for commonly-used biomaterials.

Poly(ethylene oxide) (PEO) is a synthetic polymer that has been found to be nontoxic and is approved by the United States Food and Drug Administration(FDA) as excipients and as carriers in different pharmaceutical formulations, foods, and cosmetics. [31] PEO is also soluble in water, which allows for leaching from a mixture without the use of toxic or harmful solvents. Polycaprolactone (PCL) is a semicrystalline polymer with a homopolymer or repeating structure of five nonpolar methylene groups and a polar ester group. It has been gaining increased attention in the field of bioengineering because it is biocompatible, biodegradable, and provides desirable mechanical strength. It is also approved by the FDA for some human clinical applications such as implantable devices. [32-34] PCL has similar density and melting temperature as PEO, making it ideal for creating scaffolds using PEO as the porogen. Table 1.4 shows some physical properties of PEO and PCL.

Table 1.3: Properties of commonly-used biomaterials. Adapted from Brahatheeswaran Dhandayuthapani, Yasuhiko Yoshida, ToruMaekawa, and D. Sakthi Kumar. (2011)

Table 1.4: Properties of PCL and PEO.

Density (g/cm3) Materials Tg (°C) Tm (°C) 25°C 100°C PCL -69† 60* 1.145* 1.036† PEO -67* 65* 1.13* 1.07† *Provided by Sigma Aldrich †Provided from the literature [29]

1.4.2 Ceramics By themselves, neither polymers nor ceramics are ideal for matching the properties of load- bearing tissue. Polymers tend to be very ductile and not rigid while ceramics are too stiff and brittle. Combining the two however may help overcome these limitations. Bioceramics can be divided into three categories [35]: 1. Bioinert groups (e.g., zirconia and alumina) 2. Surface bioactive groups that includes sintered HA (s-HA) 3. Bioresorbable groups that include not sintered HA (u-HA) Category 3 is generally used for biodegradable purposes to increase the strength of a structure. HA is a naturally occurring mineral form of calcium apatite that is the primary mineral component in bone. HA with incorporated growth factors has been shown to improve osteoblast cell growth. [36] HA particles were incorporated into the PEO/PCL powder mix. Previous studies have shown that the level of HA remaining in the scaffold after PEO leaching should ideally be 20% w/w for tissue engineering. This is illustrated by Figure 1.5. [37]

Prolife

Figure 1.5: Data illustrating the effect of wt.% HA on the strength of HA/polymer composites. Data were averaged and are represented with error bars in cases in which there were multiple test conditions for a given fraction of HA. Authors who performed both compressive and flexural tests are demarked by closed and open symbols, respectively. [37]

1.5 Scaffold fabrication

PCL (Mw: 70,000-90,000 Da) and PEO (Mw: 100,000 Da) were purchased from Sigma Aldrich. HA of approximate particle size of 5µm was supplied by Plasma Biotal LTD. To control pore size, only polymer and ceramic particles between 250 µm and 425 µm were used in scaffold fabrication. The provided PEO had particle sizes less than 250 µm so PEO disks were created by placing PEO into a stainless steel mold and heated at 100 °C for 25 minutes. Once the disks were cooled, they were ground in a SPEX SamplePrep 6770 Freezer/Mill using a polycarbonate grinding vial. PEO disks were ground at a rate of 12 CPD for 2 minutes after a pre-cool time of 10 minutes. PEO particles greater than 425 µm in size were collected and ground again at a rate of 8 CPS for 1 minute after a pre-cool time of 10 minutes. PCL was provided in pellet form so the pellets were ground directly at a rate of 10 CPS for 2 minutes after a pre-cool time of 10 minutes using a stainless-steel grinding vial. PCL particles greater than 425 µm were collected and ground again at a rate of 8 CPS for 1 minute after a pre-cool time of 10 minutes. PCL scaffolds were produced by mixing the appropriate amount of ground and sieved PEO and PCL using a vortex mixer. The polymer mixture was then placed in stainless-steel molds and heated for 1 hour at 100 °C. After 1 hour, the molds were removed and cooled at room temperature. The PEO/PCL disks were then placed in a bath of deionized water for 24 hours at 40°C. The HA

provided had a particle size of 5 µm so PCL/HA disks were made by mixing appropriate amounts of HA and PCL of particles size less than 250 µm using a vortex mixer. The PCL/HA mixture was then placed in stainless steel molds and heated at 100 °C for 25 minutes. The disks were then cooled at room temperature and ground using the same procedure at PEO disks. The resulting PCL/HA powder mixture was then sieved and PCL/HA scaffolds were fabricated using the same procedure as discussed above.

1.6 Porosity and interconnectivity

The effectiveness of a scaffold is greatly dependent on the pore size and interconnection between the pores. High porosity and interconnectivity throughout the scaffold allow cell growth and transport of both nutrients and metabolic waste. Large surface area to volume ratio is also desirable. Pore sizes need to be greater than the size of the cells that will seed the scaffold (approximately 100 µm for osteoblasts). [38] The recommended pore size for bone regeneration varies with different studies and with different scaffold fabrication techniques. It has been reported that bone growth on scaffolds produced by solid freeform fabrication (SFF) technique is independent of pore diameter (at a range of 300 – 800 m), with no statistical difference between pore sizes. This contrasts results using non-SFF scaffolds, where optimal pore diameters ranging from 200 µm to 600 µm have been suggested. However, unlike the single pore diameter in the designed scaffolds (SFF), non-designed scaffolds have a range of pore sizes, which may explain the different results. [9] Table 1.5 shows other recommendations for optimum pore sizes for bone tissue engineering.

Table 1.5: Optimum pore size for bone regeneration. Adapted from Sarazin, Roy, & Favis (2004)

Optimum pore size (µm) Reference 100-150 [32] 100-250 [38] 200-400 [39] 100-350 [40] 75-150 [41,42] Less than 100 [43]

As mentioned, at least a 60:40 (vol %) ratio is needed to create a co-continuous structure for melt-processed polymer blends. However, an increase in porosity leads to a decrease in mechanical strength.

1.7 Mechanical properties

The mechanical properties of scaffolds play an important role in the effectiveness of the scaffold. Generally, scaffolds used in functional tissue engineering need to match the mechanical properties of the target tissue. This allows the scaffolds to survive in the implant environment and to withstand stresses put on them by the body. The mechanical properties of various human tissues can be seen below in Table 1.6. [44]

Mechanical testing was done using an Instron 3345 Material Testing System with a 1 kN load cell. Scaffolds were tested under unconfined ramp compression at 60% strain and a displacement rate of 1 mm/min using a preload of 4.45 N. Three samples with a diameter of 5 mm and a thickness of 1-2 mm were tested for each scaffold formulation.

Table 1.6: Mechanical properties of human tissues [44]

Tensile strength Compressive Youngs’ modulus Fracture toughness (MPa) strength (MPa) (GPa) (Mpa.m1/2) Cancellous bone[45] N/a 4-12 0.02-0.5 N/a Cortical bone[45] 60-160 130-180 3-30 2-12 Cartilage[46] 3.7-10.5 N/a 0.7-15.3 (MPa) N/a Ligament[47] 13-46 N/a 0.065-0.541 N/a Tendon[47] 24-112 N/a 0.143-2.31 N/a

1.8 Cell Attachment and Growth

In addition to porosity, interconnectivity, and mechanical properties, cell attachment and growth is a critical characteristic of scaffolds used in bone tissue engineering. The ability of a scaffold to facilitate this is often predicted by assessing proliferation, gene expression, and morphology of cells seeded onto the scaffold.

In this work, mouse calvaria-derived MC3T3-E1 cells were purchased from Sigma Aldrich at passage 17. Cells were cultured in growth media composed of α-MEM supplemented with 10% fetal bovine serum and 1% penicillin/streptomycin (MEM-α + 10% FBS+ 1% Pen/Strep).

Growth media was replaced every 2-3 days and cells were incubated at 37°C in a 5% CO2 atmosphere. Subculturing was done by harvesting the cells with 0.05% trypsin/0.53 mM EDTA in Hank’s buffered salt solution (Corning). The protocol for cell passaging was as follows:

1. Discard the media of the cells 2. Wash the cells with 3 ml of trypsin/EDTA removing the trypsin/EDTA right away

3. Add 2 ml of trypsin/EDTA to each dish and put the dishes in a 37°C, CO2 incubator for ~3 min (until the cells are detached from the dish) 4. Tap the dishes to dislodge cells 5. Add 8 ml of growth media to each dish and pipette up and down (in the dish) and then transfer each dish to a 15 ml tube 6. Pellet cells by centrifugation for 5 min 7. Remove the media and suspend the cells in fresh growth media by pipetting up and down 8. Add 2 ml of growth media with cells to each dish and then add additional growth media so each dish contains a total of 12 mL of media

Scaffolds were sterilized prior to cell culture by soaking in 70% ethanol for 2 hours and then washed and left overnight in sterile phosphate buffered saline (PBS). The scaffolds were then washed with sterile PBS two more times for 1.5 hours each and then left overnight in the media used for cell culturing (MEM-α + 10% FBS+ 1% Pen/Strep). The scaffolds were then transferred into 24-well culture plates.

Cells were harvested and seeded on the scaffolds in 24-well plates at a density of 7.0 105 cells/scaffold. Cells were allowed to adhere to the scaffolds for 24 hours and then the scaffolds were transferred to new 24-well plates. This was considered day 0. Media was replaced every 2 days and samples were taken on days 0, 7, 14, and 21.

1.8.1 Scanning Electron Microscopy

Scanning electron microscopy (SEM) is a common and increasingly used technique to image the morphology of cells seeded onto a scaffold. As SEM places samples in a vacuum, biological samples must be carefully prepared for imaging. Common methods for sample preparation of cells on biomaterials can be seen in Table 1.7. [48]

Table 1.7: Comparison of different SEM sample preparation methods [48]

While cortical point drying (CPD) is generally used in preparing biological materials, there can be complications from the increase in temperature and pressure that is required during this process. HMDS is an alternative method to CPD that has been shown to be effective in preparing biological samples. [49]

The protocol for preparing scaffolds with HMDS was as follows:

1. Remove growth media

2. Wash scaffolds with PBS two times (5 minutes each) 3. Fix cells in 1 ml of 1% Glutaraldehyde + 2% Paraformaldehyde in PBS for 1 hour 4. Wash scaffolds with PBS 4 times (5 minutes each) 5. Take the scaffolds through a dehydration and HMDS embedding process i. 25% EtOH – 20 minutes ii. 50% EtOH – 20 minutes iii. 75% EtOH – 20 minutes iv. 95% EtOH – 30 minutes v. 100% EtOH – 1 hour (3 times) vi. 2:1 EtOH:HMDS – 30 minutes vii. 1:2 EtOH:HMDS – 30 minutes viii. 100% HMDS – 30 minutes (3 times) 6. Remove HMDS 7. Air dry samples overnight at room temperature

After air drying, samples were silver painted, sputter coated with Au and imaged using a Zeiss Supra 35 VP FEG scanning electron microscope.

1.8.2 Proliferation

Using a CyQUANT Cell Proliferation Assay Kit (Life Technologies), proliferations assays were carried out on samples at each time point. DNA content was measured using a modified version of the CyQuant cell proliferation assay. [50] Cell lysis buffer was made supplemented with 180 mM NaCl, 1mM EDTA, and 0.75 Kunitz/ml RNAse to ensure that only DNA content was read and that possible cell differentiation and varying RNA amounts did not interfere with the final results. Standard curves were generated with both λ DNA and MC3T3-E1 cells. To prepare cell- seeded scaffolds for proliferation, media was removed and the scaffolds were washed 2 times with PBS for 5 minutes each. The PBS was then removed and the samples stored at -80 °C until the assay was performed. The protocol for performing the proliferation assay was as follows:

1. Add 250 µL of RNAse buffer to each sample

2. Vortex 3. Sonicate for 20 seconds on ice 4. Incubate for 1 hour 5. Sonicate for 10 seconds 6. Vortex and spin down 7. Add the following amounts from the top of each sample to wells of solid black 96 well plate i. 100 µL of sample ii. 50 µL of sample with 50 µL of RNAse buffer iii. 25 µL of sample dilution with 75 µL of RNAse buffer 8. Add 100 µL of 2x dye/lysis buffer to each well i. Mix by pipetting up and down

The fluorescence of the samples were measured with a NOVOstar cell-based fast kinetic microplate reader with a 482/50 excitation filter and a 528/20 emission filter.

1.9 Project Statement

1.9.1 Hypothesis

In this study, polycaprolactone/hydroxyapatite (PCL/HA) composite scaffolds were prepared using a melt-processing/porogen-leaching technique. In an effort to produce scaffolds with optimal porosity, highly interconnected pore structure, and desired mechanical and osteoconductivity, the role of key parameters, such as polymer/porogen ratio (PCL/PEO), HA content, the presence of salt as an additional porogen, and clamping pressure were investigated. We put the emphasis on improving the interconnectivity and pore size range, so as to minimize the number of pores below 100 m. To this end, the following hypotheses were tested:

 Melt processing of PCL/PEO is an effective technique for creating scaffolds with interconnected pores, desired pore size range (> 100 m), and porosity greater than 50%. PEO acts as a porogen and an interconnected pore network is formed once it is dissolved in water. SEM imaging was used for testing this hypothesis.  Incorporation of HA can enhance the osteoconductivity without compromising the mechanical properties of the scaffolds by our technique. The produced scaffolds should possess adequate mechanical properties for bone tissue engineering. [51-54]

 The produced scaffolds, seeded with osteoblasts, can enable cell attachment and growth for in vitro bone tissue engineering (up to 3 weeks of culture).

1.9.2 Objectives

• Determine the best processing conditions, including pressure, PEO/PCL ratio, salt content, and HA content for melt processing of these scaffolds

• Determine the optimal conditions for dissolving PEO in water

• Observe the pore size and interconnectivity of the scaffolds using scanning electron microscopy

• Determine and optimize the mechanical properties of the scaffolds

• Seed the scaffolds with osteoblasts and perform biological assays to quantify cell growth and determine cell viability for a time period of up to 3 weeks

1.10 Thesis Outline

Chapter 2 presents a peer-reviewed paper published in the proceedings of the Society of Plastics Engineers ANTEC ® 2013 conference. It focuses on the optimization of scaffold mechanical properties and porosity by adjusting the polymer ratio, ceramic and salt content, and pressure applied during the fabrication process. The effect of varying each of these parameters is examined so that a factorial design of experiment (DOE) can be created to determine the optimal scaffold fabrication parameters (to maximize the scaffold modulus). This work was done under the guidance of Dr. Azizeh Yousefi and with the help of Cara Janney and Carlie Focke. Since the process parameters had a minor effect on the mechanical properties of the scaffolds, the rest of this study focused on obtaining scaffolds with adequate pore interconnectivity based on SEM analysis.

Chapter 3 presents a paper to be submitted to Bio-Medical Materials and Engineering. The focus of this paper is to estimate the overall effectiveness of two different scaffold formulations in bone tissue engineering. Polymer and polymer/ceramic scaffolds were evaluated and compared by examining the mechanical properties, porosity, and cell attachment and growth. This work

was done under the guidance of Dr. Azizeh Yousefi and Dr. Paul James. Polymer grinding, scaffold fabrication, and mechanical testing were done with the help of Cara Janney and Carlie Focke. Cell thawing, media preparation, and H&E staining (not included) were done by Rosa Akbarzadeh who also assisted with cell culture and preparation of cell seeded scaffolds for analysis. The micro-CT analysis was performed by James Schmitz at the University of Texas Health Science Center at San Antonio (UTHSCSA).

Chapter 4 provides conclusions on the results shown in Chapters 2 and 3. It discusses the significance of these findings and suggests future work.

1.11 References

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2. Coutu, D. L. et al. Hierarchical scaffold design for mesenchymal stem cell-based gene therapy of hemophilia b. Biomaterials 2011; 295-305.

3. Yousefi, A., Janssen, M., & Tan, Q. W., et al. Towards multiscale performance prediction for tissue engineering scaffolds. 2011.

4. Guilak F, Butler DL, Goldstein SA, Mooney DJ. Functional tissue engineering. Springer, New York. 2003.

5. Ingber DE, Mow VC, Butler D, Niklason L, Huard J, Mao J, Yannas I, Kaplan D, Vunjak- Novakovic G. Tissue engineering and developmental biology: going biomimetic. Tissue Engineering 2006;12:3265-3283.

6. Moutos FT, Freed LE, Guilak F. A biomimetic three-dimensional woven composite scaffold for functional tissue engineering of cartilage. Nature Mater 2007;6:162-167.

7. Yousefi AM, Gauvin C, Sun L, DiRaddo RW, Fernandes J. Design and Fabrication of 3D- Plotted Polymeric Scaffolds in Functional Tissue Engineering. Polymer Engineering and Science 2007;47:608-618.

8. Meyer, U., & Wiesmann, H. P. Bone and cartilage engineering. Germany: Springer-Verlag Berlin Hiedelberg. 2006.

9. Hollister, S. J. Porous scaffold design for tissue engineering. Nature Materials 2005; 4:518– 524.

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11. Heary RF, Schlenk RP, Sacchieri TA, Barone D, Brotea C. Persistent iliac crest donor site pain: independent outcome assessment. Neurosurgery 2002;50:510–516; discussion 516- 517.

12. Kretlow JD, Mikos AG. Review: mineralization of synthetic polymer scaffolds for bone tissue engineering. Tissue Engineering. 2007;13:927–938.

13. Nakajima T, Iizuka H, Tsutsumi S, Kayakabe M, Takagishi K. Evaluation of posterolateral spinal fusion using mesenchymal stem cells: differences with or without osteogenic differentiation. Spine 2007;32;2432-2436

14. Arrington ED, Smith WJ, Chambers HG, Bucknell AL, Davino Arrington ED, Smith WJ, Chambers HG, Bucknell AL, Davino Orthop Relat Res. 1996;329:300–309.

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16. Banwart JC, Asher MA, Hassanein RS. Iliac crest bone graft harvest donor site morbidity. A statistical evaluation. Spine. 1995;20:1055–1060.

17. Mankin HJ, Hornicek FJ, Raskin KA. Infection in massive bone allografts. Clin Orthop Relat Res. 2005;432:210–216.

18. Nishida J, Shimamura T. Methods of reconstruction for bone defect after tumor excision: a review of alternatives. Med Sci Monit. 2008;14:RA107–RA113.

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22. Hutmacher, D. W. Scaffolds in tissue engineering bone and cartilage. Biomaterials 2000;21(24): 2529-43

23. Yang, S., K. F. Leong, Z. Du and C. K. Chua. The design of scaffolds for use in tissue engineering. Part I. Traditional factors. Tissue Engineering 2001;7(6): 679-89.

24. Gomes, M. E., J. S. Godinho, D. Tchalamov, A. M. Cunha and R. L. Reis. Alternative tissue engineering scaffolds based on starch: processing methodologies, morphology, degradation and mechanical properties. Master Sci Eng 2002;C 20(1-2_: 19-26.

25. Chung, H. J., & Park, T. G. Surface engineered and drug releasing pre-fabricated scaffolds for tissue engineering. Advanced Drug Delivery Reviews 2007;59(4-5), 249-262.

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27. Shea, L. D., D. Wangm R. T. Franceschi and D. J. Mooney. Engineered bone development from a pre-osteoblast cell line on three-dimensional scaffolds. Tissue Eng 2000;6(6): 605- 17.

28. Washburn NR, Simon CG, Tona A, Elgendy HM, Karim A, Amis EJ. J Biomed Mater Res 2002;60:20–9.

29. Sarazin, P., Roy, X., & Favis, B. D. Controlled preparation and properties of porous poly( l- lactide) obtained from a co-continuous blend of two biodegradable polymers. Biomaterials 2004;25, 5965-5978.

30. Reignier, J., & Huneault, M. A. Preparation of interconnected poly (ε-caprolactone) porous scaffolds by a combination of polymer and salt particulate leaching. Polymer 2006;47, 4703-4717.

31. Fuertges, F.; Abuchowski, A. Journal of Controlled Release 1990;11, 139.

32. Maquet V, Jerome R. Design of macroporous biodegradable polymer scaffolds for cell transplantation. Mater Sci Forum 1997;250:15–42.

33. Peters MC, Mooney DJ. Synthetic extracellular matrices for cell transplantation. Mater Sci Forum 1997;250:43–52.

34. Ma PX, Zhang R. Synthetic nano-scale fibrous extracellular matrix. J Biomed Mater Res 1999;46:60–72.

35. Shikinami, Y., and Okuno, M. Bioresorbable devices made of forged composites of hydroxyapatite (HA) particles and poly-L-lactide (PLLA): Part I. Basic characteristics. Biomaterials 1999;20, 859.

36. Kim, J., Kim, I. S., & Cho, T. H, et al. Bone regeneration using hyaluronic acid-based hydrogel with bone morohogenic protein-2 and human mesenchymal stem cells. Biomaterials 2007;1830-1837.

37. Wagoner Johnson, A. J., & Herschler, B. A. A review of the mechanical behavior of cap and cap/polymer composites for applications in bone replacement and repair. Acta biomaterialia 2011;7, 16-30.

38. Whang K, Thomas CH, Healy KE, Nuber G. A novel method to fabricate bioabsorbable scaffolds. Polymer 1995;36:837–42.

39. Lu L, Mikos AG. The importance of new processing techniques in tissue engineering. MRS Bulletin, November, 1996. p. 28–32.

40. Whang K, Healy KE, Elenz DR, Nam EK, Tsai DC, Thomas CH, Nuber GW, Glorieux FH, Travers R, Sprague SM. Engineering bone regeneration with bioabsorbable scaffolds with novel microarchitecture. Tissue Eng 1999;5:35–51.

41. Li WJ, Laurencin CT, Caterson EJ, Tuan RS, Ko FK. Electrospun nanofibrous structure: a novel scaffold for tissue engineering. J Biomed Mater Res 2002;60:613–21.

42. Von Recum AF, Shannon CE, Cannon CE, Long KJ, van Kooten TG, Meyle J. Tissue Eng 1996;2:241–53.

43. It.al.a AI, Yl.anen HO, Ekholm C, Karlsson KH, Aro HT. Pore diameter of more than 100 m is not requisite for bone ingrowth in rabbits. J Biomed Mater Res (Appl Biomater) 2001;58:679–83.

44. Yang, S., Leong, K., Du, Z., & Chua, C. The design of scaffolds for use in tissue engineering. part i. traditional factors. Tissue engineering 2001;7, 679-689.

45. Yang, S.F. Study on biomimetic artificial bone [Ph.D. dissertation]. Tsinghua University, China, 1999.

46. Parsons, J.R. Cartilage. In: Black, J., and Hastings, G., eds. Handbook of Biomaterials Properties. New York: Chapman & Hall, 1998, pp. 40–46.

47. Woo, S.L.-Y., and Levine, R.E. Ligament, tendon and fascia. In: Black, J., and Hastings, G., eds. Handbook of Biomaterials Properties. New York: Chapman & Hall, 1998, pp. 59–65.

48. Lee, J. T. Y., & Chow, K. L. SEM sample preparation for cells on 3D scaffolds by freeze- drying and HMDS. Scanning 2012;34(1), 12–25. doi:10.1002/sca.20271

49. Shively, S., Miller, W. R., Shively, S., & Miller, W. R. The use of HMDS (hexamethyldisilazane ) to replace Critical Point Drying ( CPD ) in the preparation of tardigrades for SEM ( Scanning Electron Microscope ) imaging 2013;112(3), 198–200.

50. St-Pierre, J.-P., Gauthier, M., Lefebvre, L.-P., & Tabrizian, M. Three-dimensional growth of differentiating MC3T3-E1 pre-osteoblasts on porous titanium scaffolds. Biomaterials, 26(35), 7319–28. doi:10.1016/j.biomaterials 2005;05.046

51. Rezwan K, Chen QZ, Blaker JJ, Boccaccini AR. Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering. Biomaterials. 2006;27:3413–31.

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53. Hutmacher DW, Schantz JT, Lam CX, Tan KC, Lim TC. State of the art and future directions of scaffold-based bone engineering from a biomaterials perspective. J Tissue Eng Regen Med. 2007;1:245–60.

54. Martin, R.B., Burr, D.B., Sharkey, N.A. Skeletal Tissue Mechanics 1998; Springer, New York, pp. 29–30, 131–134, 143–151.

Chapter 2

Design and Fabrication of Polymer/Ceramic Scaffolds for Bone Engineering

Joshua Minton, Cara Janney, Carlie Focke, and Azizeh-Mitra Yousefi

Chemical, Paper, and Biomedical Engineering Department, Miami University, Oxford, OH

2.1 Abstract

Bone tissue engineering is a rapidly developing field, and seeks to offer an alternative treatment for bone defects by restoring and maintaining the function of bone tissue. One of the most established approaches is using polymer scaffolds seeded with osteoblast and other growth factors to speed the body’s natural healing processes, decreasing rehabilitation time for patients. The biomimetic design of the scaffolds will need to replicate the structural and mechanical properties of the tissue and be stiff enough to withstand immediate weight bearing. The effectiveness of this approach is determined by examining the properties of the scaffold including porosity, interconnectivity, and mechanical properties. The goal of this study is to create viable polymer/ceramic scaffolds through melt processing of polycaprolactone (PCL) and poly(ethylene oxide) (PEO), combined with hydroxyapatite (HA) and salt (NaCl), followed by porogen leaching. The effects of polymer ratio, ceramic and salt content, and the pressure applied during the fabrication process have been examined in this study. These results will be used to create a factorial design of experiments (DOE) to determine the optimal scaffold fabrication parameters.

2.2 Introduction

greatly decreases the effectiveness of the treatment. One solution to this problem is to use polymer scaffolds that have been found to improve cell viability in vivo. [1] Polymer scaffolds have also proven useful in gene therapy, a technique for treating both acquired and hereditary diseases. Gene therapy has shown great promise in treating hemophilia B and other hereditary plasma protein deficiencies; however, the same problem of survival rate in vivo persists. [2] A potential solution to this is the use of polymer scaffolds which have been shown to support long term cell growth and protein delivery in vivo and have been used to cure hemophilic mice for over 12 weeks. [2,3]

The focus of this study is on the use of polymer/ceramic scaffolds in bone tissue engineering. The rapid restoration of tissue biomechanical function remains a challenge, and emphasizes the need to replicate the structural and mechanical properties of the tissue using new scaffold designs. [4] To address this challenge, the focus of the current efforts is on designing biomimetic scaffolds. [5,6] Producing scaffolds with bone-mimicking mechanical properties has many potential advantages upon in vivo implantation. The stiffness would be adequate to withstand immediate weight bearing. Moreover, it would allow faster rehabilitation for patients.

The effectiveness of tissue-engineering scaffolds is affected by porosity, interconnectivity of the pores, and mechanical properties of the chosen scaffold material. Porosity and pore interconnectivity throughout the scaffold are required to promote cell loading and migration, tissue and vasculature growth, and adequate transport of oxygen and nutrients. A major problem with traditional methods of fabricating these scaffolds is that they result in a random pore structure, leading to poor permeability and interconnectivity. [7]

The goal of this study is to develop a method of fabricating polymer scaffolds that will lead to controlled pore sizes and fully interconnected pores, while avoiding the use of toxic solvents. Our scaffolds are produced through melt processing, which involves heating a blend of two polymer powders above their melting temperatures. Polycaprolactone (PCL) is the primary scaffold material used in this study due to its low melting temperature (below 70°C). Poly(ethylene oxide) (PEO) has a similar melting point and is used as the primary porogen, along with salt particles. The porosity and interconnectivity of the scaffolds are achieved through dissolving PEO and salt in water. In addition, in bone tissue engineering there is a need for osteoconductive scaffold materials. Incorporation of ceramic particles, e.g. hydroxyapatite (HA),

can result in materials with the desired mechanical properties and better biocompatibility and osteoconductivity.

In an effort to produce scaffolds with optimal porosity, highly interconnected pore structure, desired mechanical properties, and osteoconductivity, the role of key parameters, such as polymer/porogen ratio (PCL/PEO), HA content, additional porogen content (NaCl), and pressure applied during the fabrication process, have been investigated. We have put the emphasis on improving the interconnectivity and pore size range so as to minimize the number of pores below 100 µm, while maintaining adequate mechanical properties. To this end, we aim to test the following hypotheses:

 Melt processing of PCL/PEO is an effective technique for creating scaffolds with interconnected pores, desired pore size range (> 100 µm), and porosity greater than 50%. PEO acts as a porogen and an interconnected pore network is formed once it is dissolved in water. Scanning Electron Microscopy (SEM) and permeability measurements will provide a quantitative measure for testing this hypothesis.  Incorporation of HA will enhance the osteoconductivity without compromising the mechanical properties of the scaffolds by our technique. The produced scaffolds should possess adequate mechanical properties for bone tissue engineering.

2.3 Scaffold Materials

Polymers have been widely used as biomaterials for the fabrication of medical device and tissue- engineering scaffolds. Polycaprolactone (PCL) is a semicrystalline polymer with a repeating structure of five nonpolar methylene groups and a polar ester group. It has been gaining increased attention in the field of bioengineering because it is biocompatible, biodegradable, and provides desirable mechanical strength. It is also approved by the United States Food and Drug Administration (FDA) for some human clinical trials such as implantable devices. [8-10] The PCL homopolymer has a degradation time of 2 years and PCL structures have been found to be morphologically stable for about 1 year. [11] Poly(ethylene oxide) (PEO) is a synthetic polymer that has been found to be nontoxic and is approved by the FDA as excipients and as carriers in different pharmaceutical formulations, foods, and cosmetics. [12] PEO is soluble in water, and

therefore can be leached out from a mixture without the use of toxic or harmful solvents. PEO also has similar density and melting temperature as PCL, making it ideal as the porogen for creating scaffolds. Table 1 shows some physical properties of PCL and PEO.

Table 2.1: Properties of PCL and PEO.

Density (g/cm3) Materials Tg (°C) Tm (°C) 25°C 100°C PCL 69† 60* 1.145* 1.036† PEO 67* 65* 1.13* 1.07† *Provided by Sigma Aldrich

†From the literature [13]

By themselves, neither polymers nor ceramics are ideal for matching the properties of load- bearing tissues. Polymers tend to be very ductile and not rigid while ceramics are too stiff and brittle. Combining the two, however, may help overcome these limitations. Bioceramics can be divided into three categories [14]:

1. Bioinert groups such as alumina and zirconia 2. Surface bioactive groups such as sintered HA, bioglass, alumina-wollastonite glass ceramic 3. Bioresorbable groups such as non-sintered HA, a- or b-tricalcium phosphate, tetracalcium phosphate, and octacalcium phosphate

Category 3 is generally used for biodegradable purposes to increase the strength of a structure. The mechanical properties of scaffolds play an important role in the effectiveness of the scaffold. Generally, scaffolds used in functional tissue engineering need to match the mechanical properties of the target tissue. The mechanical properties of load-bearing human tissues are listed in Table 2. [15] HA is a naturally occurring mineral form of calcium apatite that is the primary mineral component in bone. HA with incorporated growth factors has been shown to improve osteoblast cell growth. [16]

Table 2.2: Mechanical properties of human tissues [15]

Tensile strength Compressive strength Young’s modulus (MPa) (MPa) (GPa) Cancellous N/a 4-12 0.02-0.5 bone[17] Cortical bone[17] 60-160 130-180 3-30 Cartilage[18] 3.7-10.5 N/a 0.7-15.3 (MPa) Ligament[19] 13-46 N/a 0.065-0.541 Tendon[19] 24-112 N/a 0.143-2.31

2.4 Materials

PCL (Mw: 70,000-90,000) and PEO (Mw:100,000) were purchased from Sigma-Aldrich. Scaffolds with PCL/PEO compositions of 50:50, 40:60, 35:65, and 30:70 (wt %) were produced by compression-molding at 100°C. Scaffold pore size was controlled by limiting the size of the polymer particles to between 250 µm and 425 µm prior to melt processing. This was achieved by grinding and sieving of both polymer particles. HA particles were also incorporated into the PEO/PCL powder mix. The particle size of the HA was ~ 5µm, and was kindly supplied by Plasma Biotal LTD. The level of HA remaining in the scaffold after PEO leaching should ideally be 20% w/w for bone tissue engineering. [20] To produce PCL/HA scaffolds, HA particles were blended with PCL powder using a vortex mixer and compression molded. Scaffolds with three different HA concentrations were fabricated: 10%, 20%, and 30% w/w. To investigate the effect of additional porogen on pore interconnectivity, NaCl particles were also used in this study. Various particles sizes of NaCl were examined to determine the effect of particle size on the mechanical properties and scaffold architecture. The produced constructs were placed in a bath of deionized water at 40°C for 24 hours to leach out the PEO and NaCl.

2.5 Scaffold Characterization

The scaffolds were tested under unconfined ramp compression using Instron 3345 with a 1 kN load cell, at 60% strain and a displacement rate of 1 mm/min. The compressive Young’s modulus (linear slope of the stress-strain curve) as well as the modulus at 10% strain was estimated for the scaffolds. These data helped to eliminate the fragile constructs and to identify the appropriate designs for the future in vitro trials. The scaffolds were imaged using SEM to inspect their morphology and pore interconnectivity. The scaffold porosity was calculated using the equation as follows:

Va  Vt φ  100 Va

where Va is the apparent volume of the scaffold estimated based on the geometry of each disk

(thickness and diameter), and Vt is the true volume of each scaffold calculated based on the combined polymer/HA matrix density (rule of mixtures) and scaffold mass (m) using . The scaffold density was calculated using .[7] Vt  m/ matrix   matrix(1)

2.6 Statistical Analysis

The student’s T-test was used to analyze the statistical significance of differences between the different samples. Comparisons were made for two groups at a time, and the p values < 0.05 were considered significant.

2.7 Results and Discussion 2.7.1 Effect of Polymer Ratio

The results of mechanical testing and porosity calculations for scaffolds with various polymer ratios can be seen in Figure 2.1 and 2.2. These results indicate that a higher ratio of PEO to PCL

leads to decreased mechanical strength and increased porosity, with porosity being approximately equal to the weight percent of PEO.

Young's Modulus 2.5

2.0

1.5

1.0 Modulus (MPa) Modulus

0.5

0.0 PEO/PCL 60/40 PEO/PCL 65/35 PEO/PCL 70/30

Figure 2.1: Average Young’s modulus for scaffolds with various PEO/PCL polymer ratios.

Porosity 75

Porosity (%) Porosity 60

PEO/PCL 60/40 PEO/PCL 65/35 PEO/PCL 70/30

Figure 2.2: Average calculated porosity for scaffolds with various polymer ratios

Three ranges of salt particle size were examined to determine the effects on mechanical properties and porosity: less than 100 µm, between 100 µm and 250 µm and 250 µm and 425

µm. Salt was added at 20% and 40% w/w. The results of the mechanical tests and porosity calculations can be seen in Figure 2.3 and 2.4. As Figure 2.3 shows, there is an increase in mechanical properties in scaffolds with 20% salt with particles sizes less than 100 µm as well as scaffolds with 40% salt with particle sizes between 100 µm and 250 µm. T-tests were performed to test the significance of the changes in mechanical properties, and it was found that none of the salt formulations resulted in significant increase or decrease in mechanical properties. These results can be seen in Table 2.3. To determine if the presence of residual salt particles could be the cause of the variations in mechanical properties, SEM images were taken of leached scaffolds. The results can be seen in Figure 2.5 and 2.6.

Young's Modulus 3.0

2.5

2.0

1.5

Modulus (MPa) Modulus 1.0

0.5

0.0 PEO/PCL 60/40 PEO/PCL 60/40 20% Salt <100 PEO/PCL 60/40 20% Salt 100-250 PEO/PCL 60/40 40% salt 100-250 PEO/PCL 60/40 40% salt 250-425

Figure 2.3: Average Young’s modulus for the scaffolds with various salt particles sizes and concentrations

Porosity 75

Porosity (%) Porosity 60

50 PEO/PCL 60/40 PEO/PCL 60/40 20% Salt<100 PEO/PCL 60/40 20% Salt 100-250 PEO/PCL 60/40 40% salt 100-250 PEO/PCL 60/40 40% salt 250-425

Figure 2.4: Average calculated porosity for the scaffolds with various salt particles sizes and concentrations

Table 2.3: T-test results comparing mechanical properties of the various salt concentrations and particle sizes tested. Numbers greater than 0.05 indicate that the difference in mechanical properties is not statistcally significant.

PEO/PCL PEO/PCL 60/40 20% 60/40 salt 100 µm-250 µm PEO/PCL 60/40 - 0.373 PEO/PCL 60/40 20% salt <100 µm 0.149 -

PEO/PCL 60/40 20% salt 100 µm-250 µm 0.373 -

PEO/PCL 60/40 40% salt 100 µm-250 µm 0.123 0.110

PEO/PCL 60/40 40% salt 250 µm-425 µm 0.096 -

SEM images revealed that residual NaCl particles were present in the scaffolds with 20% salt particles less than 100 µm in size, which could be the cause of the observed increase in mechanical properties. SEM images of the scaffolds with 40% salt particles between 150 µm and 250 µm in size did not show any residual salt particles, which was confirmed by cross sectional images (not shown).

Figure 2.5: SEM image of a PEO/PCL 60/40 scaffold Figure 2.6: SEM image of PEO/PCL scaffold with 40% with 20% salt with particle sizes less than 100 µm. The salt with particle sizes between 150 µm and 250 µm. No image shows that residual salt is present in the scaffold salt particles appear to be present in the scaffold after after leaching. leaching.

2.7.2 Effect of HA concentration

The results of the mechanical tests and porosity calculations for PCL/HA scaffolds can be seen in Figures 2.7 and 2.8. Scaffolds with 20% HA content showed the highest compression modulus, which confirms the findings reported in the literature. T-tests were performed to determine if the addition of HA to the scaffold had a statistically significant effect on the mechanical properties. The results are shown in Table 2.4. SEM images of PEO/PCL 60/40 scaffolds and PEO/PCL 60/40 scaffolds with 20% HA revealed that the addition of HA to the scaffold formulation did not interfere with scaffold morphology or pore formation. This is illustrated in Figure 2.9 and 2.10.

Young's Modulus 2.5

2.0

1.5

1.0

Modulus (MPa) Modulus 0.5

0.0 PEO/PCL 60/40 10% HA PEO/PCL 60/40 20% HA PEO/PCL 60/40 30% HA

Figure 2.7: Average Young’s modulus for the scaffolds with various HA concentrations

Porosity 75

65 Porosity (%) Porosity 60

50 PEO/PCL 60/40 10% HA PEO/PCL 60/40 20% HA PEO/PCL 60/40 30% HA

Figure 2.8: Average calculated porosity for the scaffolds with various HA concentrations

Table 2.4: T-test results comparing the mechanical properties of the scaffolds with different HA concentrations. Numbers less than 0.05 (bold) indicate that the difference is statistcally significant. PEO/PCL PEO/PCL 60/40 PEO/PCL 60/40 PEO/PCL 60/40 40/40 10% HA 20% HA 30% HA PEO/PCL 40/40 - 0.017 0.446 0.012 PEO/PCL 60/40 0.017 - 0.058 0.190 10% HA

PEO/PCL 60/40 0.446 0.058 - 0.082 20% HA PEO/PCL 60/40 0.012 0.190 0.082 - 30% HA

1 mm 1 mm

Figure 2.9: SEM image of PEO/PCL 60/40 scaffold. Figure 2.10: SEM image of PEO/PCL 60/40 scaffold with 20% HA

2.7.3 Effect of Pressure

PEO/PCL 60/40 scaffolds and PEO/PCL 60/40 scaffolds with 10% HA were fabricated both with applied pressure and without. The results of the mechanical testing and porosity calculations are shown in Figure 2.11 and 2.12. In both cases, scaffolds fabricated with applied pressure showed a higher compression modulus than similar scaffolds fabricated without pressure. T-test values of 0.31 for PEO/PCL 60/40 scaffolds and 0.07 for PEO/PCL 60/40 scaffolds with 10% HA indicate that the difference is not statistically significant.

Figure 2.11: Average Young’s modulus for the scaffolds with various HA concentrations and pressure level

Figure 2.12: Average calculated porosity for the scaffolds with various HA concentrations and pressure level

2.8 Conclusions

The results of this study confirm that melt processing is an effective method for creating scaffolds with interconnected pores, desired pore size range (> 100 µm), and porosity greater than 50%. Also, the incorporation of HA into the scaffolds was successful and did not negatively

impact scaffold morphology or pore formation. Mechanical properties and porosity of the final scaffolds could be affected by several factors including HA content, polymer ratio, NaCl concentration and particle size, and pressure applied during the fabrication process. The results of this study will be used to create factorial design of experiments (DOE) and response-surface analysis. The design outcome is to optimize the mechanical properties while maintaining adequate pore interconnectivity, which will be verified by permeability measurements.

2.9 References

1. Blimenthal, B., Golsong, P., Poppe, A., Heilmann, C., Schlensak, C., Beyersdorf, F., & Siepe, M. (2010). Polyurethane scaffolds seeded with genetically engineered skeletal myoblasts: a promising tool to regenerate myocardial function. Artificial Organs, 34(2), E46-E54.

2. Coutu, D. L. et al. (2011). Hierarchical scaffold design for mesenchymal stem cell-based gene therapy of hemophilia b. Biomaterials, 295-305.

3. Yousefi, A., Janssen, M., & Tan, Q. W., et al. Towards multiscale performance prediction for tissue engineering scaffolds.

4. Guilak F, Butler DL, Goldstein SA, Mooney DJ. Functional tissue engineering (2003). Springer, New York.

5. Ingber DE, Mow VC, Butler D, Niklason L, Huard J, Mao J, Yannas I, Kaplan D, Vunjak- Novakovic G. Tissue engineering and developmental biology: going biomimetic. TISSUE ENGINEERING (2006). 12:3265-3283

6. Moutos FT, Freed LE, Guilak F. A biomimetic three-dimensional woven composite scaffold for functional tissue engineering of cartilage. Nature Mater (2007). 6:162-167.

7. Yousefi AM, Gauvin C, Sun L, DiRaddo RW, Fernandes J (2007). “Design and Fabrication of 3D-Plotted Polymeric Scaffolds in Functional Tissue Engineering”, Polymer Engineering and Science, 47:608-618.

8. Maquet V, Jerome R. Design of macroporous biodegradable polymer scaffolds for cell transplantation. Mater Sci Forum 1997;250:15–42.

9. Peters MC, Mooney DJ. Synthetic extracellular matrices for cell transplantation. Mater Sci Forum 1997;250:43–52.

10. Ma PX, Zhang R. Synthetic nano-scale fibrous extracellular matrix. J Biomed Mater Res 1999;46:60–72.

11. Barbanti, S. H., Santos Jr., A. R., Zavaglia, C. A. C., & Duck, E. A. R. (2011). Poly(e- caprolactone) and poly(d,l-lactic acid-co-glycolic acid) scaffolds used in bone tissue engineering prepared by melt compression–particulate leaching method. J. Mater Sci: Mater Med, 22, 2377-2385.

12. Fuertges, F.; Abuchowski, A. Journal of Controlled Release, 11, 139 (1990).

13. Sarazin, P., Roy, X., & Favis, B. D. (2004). Controlled preparation and properties of porous poly(l-lactide) obtained from a co-continuous blend of two biodegradable polymers. Biomaterials, 25, 5965-5978.

14. Shikinami, Y., and Okuno, M. Bioresorbable devices made of forged composites of hydroxyapatite (HA) particles and poly-L-lactide (PLLA): Part I. Basic characteristics. Biomaterials 20, 859, 1999.

15. Yang, S., Leong, K., Du, Z., & Chua, C. (2001). The design of scaffolds for use in tissue engineering. part i. traditional factors. Tissue engineering, 7, 679-689. 16. Kim, J., Kim, I. S., & Cho, T. H, et al. (2007). Bone regeneration using hyaluronic acid- based hydrogel with bone morohogenic protein-2 and human mesenchymal stem cells. Biomaterials, 1830-1837.

17. Yang, S.F. Study on biomimetic artificial bone [Ph.D. dissertation]. Tsinghua University, China, 1999.

18. Parsons, J.R. Cartilage. In: Black, J., and Hastings, G., eds. Handbook of Biomaterials Properties. New York: Chapman & Hall, 1998, pp. 40–46.

19. Woo, S.L.-Y., and Levine, R.E. Ligament, tendon and fascia. In: Black, J., and Hastings, G., eds. Handbook of Biomaterials Properties. New York: Chapman & Hall, 1998, pp. 59–65.

20. Wagoner Johnson, A. J., & Herschler, B. A. (2011). A review of the mechanical behavior of cap and cap/polymer composites for applications in bone replacement and repair. Acta biomaterialia, 7, 16-30.

Chapter 3

Polymer/Ceramic Scaffolds for Bone Engineering: Fabrication, Analysis, and Cell Growth

Joshua Mintona, Cara Janneya, Rosa Akbarzadeha, Carlie Fockea, James Schmitzb,

Paul Jamesc, and Azizeh-Mitra Yousefia

a Department of Chemical, Paper, and Biomedical Engineering, Miami University, Oxford, OH

b Center for Bone and Mineral Imaging, University of Texas Health Science Center at San Antonio, TX

c Department of Biology, Miami University, Oxford, OH

3.1 Abstract

This study examines the potential use of porous polycaprolactone and polycaprolocatone/ hydroxyapatite scaffolds fabricated through melt processing and porogen leaching for use in bone tissue engineering. Poly(ethylene oxide) was chose as a porogen due to having similar density and melting point as polycaprolactone. Pore size of the scaffold was controlled by limiting the size of polymer and porogen particles used in fabrication. Mechanical testing was used to compare the modulus of the scaffolds to that of bone tissue. The porosity and pore interconnectivity were examined with microcomputed tomography. Scanning electron microscopy was used to examine the effect on scaffold morphology caused by the addition of hydroxyapatite. Mouse calvaria-derived MC3T3-E1 cells were used to determine whether cells could attach and proliferate on scaffolds and scanning electron microscopy was used to corroborate these results. Scanning electron microscopy was also used to qualitatively compare PCL and PCL/HA scaffolds seeded with MC3T3-E1 cells. For polycaprolactone scaffolds, DNA content increased until a maximum was reached at day 14. A decrease in DNA content was seen at day 21 compared to day 14. Scanning electron microscopy images revealed that hydroxyapatite could be incorporated into polycaprolactone scaffolds without negatively effecting scaffold morphology or pore formations. SEM images also revealed a significant

increase in cell number and attachment with the incorporation of hydroxyapatite into the scaffolds.

3.2 Introduction

Bone tissue engineering is a growing field that has the potential to impact the lives of millions of people by improving the traditional methods used to treat bone defects. From 1998 to 2005, the number of joint arthroplasties and revision surgeries increased from 700,000 to over 1.1 million in the US. [1] In 2003, it was estimated that over $20 billion of medical expenses was related to fracture, reattachment, and replacement of hip and knee joints. This was predicted to increase to over $74 billion by 2015. [1,2] Bone is a specialized connective tissue that provides the body with both mechanical support and protection. It is able to undergo some level of regeneration when damage occurs from trauma, tumors, or bone disease depending on several factors including the size of the defect, nutrition, and vascularization. Several treatment options are available when the body is not able to heal the damaged bone by itself. One of these options is autogenous bone grafting which involves filling in the defect with bone tissue taken from another location in the body. Although this is the current gold standard, it comes with a risk of complications including donor site morbidity, pain, prolonged hospitalization and rehabilitation, increased risk of deep infection, hematoma, and inflammation. [3-9] An alternative option is allograft bone grafting, which uses bone tissue from other humans to fill in the defect. Possible complications of this procedure include donor-to-recipient infection, disease transmission, and host immune response. [10,11] Bone tissue engineering seeks to improve upon these limitations.

Tissue engineering involves transplanting a biofactor (such as cells) within a porous degradable scaffold. The scaffold needs to match the physical properties and mechanical function of the tissue being replaced while having a porous structure that facilitates mass transport and aids in biological delivery and tissue regeneration. [12] Such scaffolds aide the body’s natural healing processes, giving them the potential to overcome the challenges that come with traditional repair methods. The effectiveness of scaffolds for tissue engineering is determined by a number of factors including mechanical properties, pore size, pore interconnectivity, cell growth, and cell attachment.

In functional tissue engineering, scaffolds need to match the mechanical properties of the target tissue. This allows the scaffolds to withstand the natural stresses put on them by the body and prevents the inner pore structure from collapsing. The mechanical properties of various human tissues are shown in Table 3.1. [13]

Table 3.1: Mechanical properties of human tissues [13]

Tensile strength Compressive Youngs’ modulus (MPa) strength (MPa) (GPa) Cancellous bone[14] N/A 4-12 0.02-0.5 Cortical bone[14] 60-160 130-180 3-30 Cartilage[15] 3.7-10.5 N/A 0.7-15.3 (MPa) Ligament[16] 13-46 N/A 0.065-0.541 Tendon[16] 24-112 N/A 0.143-2.31

Porosity, pore size, and pore interconnectivity affect the function of the scaffold by allowing for cell loading and migration as well as the transport of nutrients and waste. A major obstacle in tissue engineering is that many traditional fabrication methods result in scaffolds with random pore structure, leading to poor interconnectivity. [17] Pore sizes also need to be greater than the size of the cell the scaffold is seeded with and the pores need to be continuous throughout the scaffold. For osteoblasts, pore sizes need to be greater than 100 µm. [18]

The method used to fabricate scaffolds has a critical effect on microstructure and pore morphology of the final scaffold. Melt processing is a technique that involves the use of two immiscible polymers. These polymers are mixed together and heated above the melting point of both polymers causing them to fuse together creating a continuous, interconnected structure. [19,20] Leaching one of these polymers from the structure using a solvent creates the final interconnected pore network. Final pore size is controlled by limiting the size of the polymer and porogen particles used in this process.

Polycaprolactone (PCL) is a semicrystalline polymer that biocompatible and biodegradable. It is also been approved for some human clinical uses such as implantable devices. [21-23] The PCL homopolymer has a degradation time of 2 years and PCL structures have been found to be

morphologically stable for around 1 year. [24] Poly(ethylene oxide) (PEO) is a synthetic, nontoxic polymer that has been used as excipients and as carriers in pharmaceutical formulations, foods, and cosmetics by the FDA. PEO is soluble in water, allowing it to be leached from a scaffold without the use of toxic solvents and it has a similar density and melting temperature as PCL. By themselves however, polymers are not ideal for matching the properties of load bearing tissues, often proving to be ductile and not rigid. Alone, ceramics are often too stiff and brittle but mixing polymer and ceramics together can overcome the limitations of both. [25] Hydroxyapatite (HA) is a naturally occurring mineral form of calcium apatite which is the primary mineral component of bone. HA with incorporated growth factors has been shown to improve osteoblast cell growth. [26] For tissue engineering the ideal concentration of HA in the final scaffold is 20% w/w. [27]

In this study, PCL and PCL/HA composite scaffolds were prepared using a melt- processing/porogen-leaching technique. The scaffolds were seeded with MC3T3-E1 osteoblastic cells in cultured medium and cell proliferation was evaluated in vitro for 21 days.

3.3 Materials and Methods 3.3.1 Materials

PCL (Mw: 70,000-90,000) and PEO (Mw: 100,000) were purchased from Sigma-Aldrich. Some properties of these two polymers are listed in Table 3.2. HA with an approximate particle size of 5µm was supplied by Plasma Biotal LTD. MC3T3-E1 osteoblastic cells were purchased from Sigma Aldrich.

Table 3.2: Properties of PCL and PEO

3 Tg Tm Density (g/cm ) Materials (°C) (°C) 25°C 100°C PCL 69† 60* 1.145* 1.036† PEO 67* 65* 1.13* 1.07† *Provided by Sigma Aldrich

†From the literature [20]

3.3.2 Fabrications of PCL and PCL/HA scaffolds

Only particles between 250 µm and 425 µm were used in scaffold fabrication to control final pore size. The provided PEO was less than 250 µm in size so PEO disks were fabricated by placing PEO into a stainless steel mold. The molds were then heated at 100 °C for 25 minutes, removed, and cooled at room temperature. Once the disks were cooled, they were ground in a SPEX SamplePrep 6770 Freezer/Mill and sieved. PCL was provided in pellet form so the pellets were ground directly and sieved. PCL scaffolds were produced by mixing PEO and PCL particles between 250 µm and 425 µm in size using a vortex mixer, placed in stainless-steel molds, and heated for 1 hour at 100 °C. The molds were then removed and cooled at room temperature. The resulting PEO/PCL disks were placed in a bath of deionized water at 40°C for 24 hours to completely leach out PEO. The HA provided had a particle size of 5 µm so PCL/HA disks were made by mixing HA and PCL of particles size less than 250 µm using a vortex mixer. The PCL/HA mixture was then heated at 100 °C for 25 minutes in stainless steel molds. The disks were cooled at room temperature, ground, and sieved. PCL/HA scaffolds were fabricated using the same procedure as discussed above. PCL scaffolds were produced at a composition of 35:65 PCL/PEO and PCL/HA scaffolds were produced with the same composition. These compositions were chosen based on the results of our preliminary studies using different PCL/PEO and PCL/HA ratios. [28] HA concentration after porogen leaching should be ideally 20% weight per weight HA to PCL. [27]

3.3.3 Characterization of porous scaffolds

Scaffolds were tested under unconfined ramp compression using an Instron 3345 Material Testing System with a 1kN load cell. Compression tests to 60% strain were carried out using a displacement rate of 1mm/min after a pre-load of 4.45 N. Three samples (n=3) with a diameter of 5 mm and a thickness of 2 mm were tested for each scaffold formulation. The Young’s modulus and the modulus at 10% strain were calculated from the stress-strain curve.

Scanning electron microscopy (SEM) was used to visualize scaffold morphology and pore formation. In addition, since HA (Ca10(PO4)6(OH)2) contains P and Ca which is not present in PCL or PEO, energy-dispersive X-ray spectroscopy (EDS) and elemental mapping were used to identify the presence of HA in PCL/HA scaffolds and to determine its dispersion in the final

structure. Thermogravimetric analysis (TGA) was used to quantitatively verify the final concentration of HA is PCL/HA by determining mass loss in air. Porosity and morphology information was gathered using microCT analysis performed at the University of Texas Health Science Center at San Antonio.

3.3.4 Cell culture

Mouse calvaria-derived MC3T3-E1 cells were cultured in α-MEM medium supplemented with 10% fetal bovine serum and 1% penicillin/streptomycin (MEM-α + 10% FBS+ 1% Pen/Strep) at 37°C in an atmosphere with 5% CO2. Media was replaced every 2-3 days and cells were subcultured through trypsinization.

Scaffolds were sterilized prior to cell culture by soaking in 70% ethanol for 2 hours and then washed and left overnight in PBS. The scaffolds were then washed with PBS two more times for 1.5 hours each and then left overnight in the media used for cell culturing (MEM-α + 10% FBS+ 1% Pen/Strep). The scaffolds were then transferred into 24-well culture plates. Cells were harvested and seeded on the scaffolds in 24-well plates at a density of 7 × 105 cells/scaffold. Cells were allowed to adhere to the scaffolds for 24 hours and then the scaffolds were transferred to new 24-well plates. This was considered day 0. Medium was replaced every 2 days and samples were taken on days 0, 7, 14, and 21.

3.3.5 Proliferation assay

DNA content at each time point was measured using a modified version of the CyQUANT cell proliferation assay (Life Technologies). [28] Cell-seeded scaffolds were washed twice with PBS and stored at -80 °C until the assay was performed. Scaffolds were thawed at room temperature and 250 µL of 1x CyQUANT cell-lysis buffer supplemented with 180 mM NaCl, 1mM EDTA, and 0.75 Kunitz/ml RNAse was added to achieve cell lysis. Samples were sonicated followed by a 1 hour incubation at room temperature. After incubation, samples were sonicated a second time and spun down. Then, 100 µL of the cell lysate was mixed with 100 µL of 2x CyQUANT GR dye in lysis buffer in solid black, 96 well microplates. The fluorescence of the samples were measured with a NOVOstar cell-based fast kinetic microplate reader with a 482/50 excitation

filter and a 528/20 emission filter. Standard curves were generated using λ DNA and known quantities of MC3T3-E1 cells.

3.3.6 Preparation for SEM imaging

PCL and PCL/HA scaffolds were imaged and compared after being air dried for 24 hours then silver painted and sputter-coated with Au. Cell-seeded scaffolds to be analyzed with SEM were removed from the media, washed twice in PBS, and fixed with 1% glutaraldehyde + 2% paraformaldehyde in PBS for 1 hour. The scaffolds were then put through serial dehydration in ethanol. In place of Critical Point Drying (CPD), scaffolds were prepared with hexamethyldisilazane (HMDS). After the final immersion in 100% ethanol, scaffolds were immersed in a solution of 2:1 ethanol:HMDS for 30 minutes followed by a solution of 1:2 ethanol:HMDS for 30 minutes. Finally, the scaffolds were immersed in 100% HMDS 3 separate times for 30 minutes each. The scaffolds were air dried for 24 hours then silver painted and sputter-coated with Au. Images were taken using a Zeiss Supra 35 VP FEG scanning electron microscope.

3.3.7 Statistical Analysis

All quantitative data are expressed as mean  standard deviation (n=3). The student’s t-test was used to analyze the statistical significance of differences between the samples. Comparisons were made for two samples at a time, and the p values < 0.05 were considered significant.

3.4 Results and Discussion 3.4.1 Characterization of porous scaffolds

SEM images showed that HA could be incorporated into the scaffold without negatively effecting scaffold morphology or pore formation. This is shown in Figure 3.1 and 3.2. EDS confirmed the presence of P and Ca in the final PCL/HA scaffold and elemental mapping was used to identify the location of each element in the final scaffold (Figure 3.3). By overlaying these images, it can be shown that HA is dispersed throughout the surface of the scaffold, not present in only isolated sections. This is shown in Figure 3.4. TGA analysis of scaffolds

fabricated with 20% HA determined a final concentration of HA of to be 17.66% when evaluated in air. The difference between the fabricated concentration of 20% and the final calculated concentration of 17.66% could be the result of HA being removed during porgen leaching or of human error during the fabrication process. These results are shown in Figure 3.5.

Figure 3.1: SEM image of PCL scaffold Figure 3.2: SEM image of PCL/HA scaffold

Figure 3.3: Elemental mapping images of O (yellow), P Figure 3.4: Elemental mapping images overlaid onto a (teal), and Ca (purple) in PCL/HA scaffold. single image to show the distribution of HA in PCL/HA scaffold.

Figure 3.5: Thermogravimetric analysis of PCL/HA scaffolds fabricated with 20% HA

Mechanical testing also revealed that the incorporation of HA into the scaffold increased the modulus at 10% strain (p<0.05), while the effect on the Young’s modulus was not statistically significant. A typical stress-stress curve obtained can be seen in Figure 3.6. Table 3.3 shows the results obtained from both the mechanical testing and porosity calculations. Both Young’s modulus and the modulus at 10% strain are lower than the reported modulus for bone tissue. [13] Therefore, these constructs would be suitable as bone graft substitutes for non-load bearing applications.

4 Sample 1

3 Sample 2 Sample 3

Stress (MPa) Stress 1

0 0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 Strain

Figure 3.6: Stress-strain curve of PCL/HA scaffolds

Table 3.3: Results of physical characterization of scaffolds

Young’s Modulus (MPa) Modulus at 10% strain (MPa) Porosity (%)

PCL scaffolds 1.41 ± 0.28 1.53 ± 0.23 68.5 ± 0.99 PCL/HA scaffolds 1.34 ± 0.57 2.13 ± 0.31 65.7 ± 0.65

MicroCT analysis of PCL and PCL/HA scaffolds determined a higher porosity than the calculated porosity as well as the percent of open pores and percent of closed pores. For both PCL and PCL/HA scaffolds, the percent of closed porosity is less than 0.5% indicating that almost all of the pores formed in the final scaffold are open and would facilitate the mass transport. These results can be seen in Table 3.4. Additionally, microCT analysis was able to determine the distribution of pore sizes throughout the entire scaffold (Figure 3.7) which showed both PCL and PCL/HA scaffolds having over 90% of pores greater than 106 µm in size. A summary of these results can be seen in Table 3.5.

Table 3.4: Porosity information for PCL and PCL/HA scaffolds determined by microCT analysis

Porosity (%) Open Porosity (%) Closed Porosity (%)

PCL scaffolds 82.1 82.0 0.45 PCL/HA scaffolds 76.8 76.7 0.43

Table 3.5: Pore size distribution of PCL and PCL/HA scaffolds

Pores > 106 µm (%) Pores 247-428 µm (%)

PCL scaffolds 96.3 33.6 PCL/HA scaffolds 92.8 37.3

Figure 3.7: Pore size distribution of PCL and PCL/HA scaffolds (where separation range indicates pore size and scaffold separation distribution represents the percent of total scaffold with the indicated pore size)

Typical melt processing techniques rely on the use of an extruder to mix the polymer blends before melting and often result in scaffolds with pore sizes smaller than 100 µm. [17,30,31] The results of this study show that by eliminating the use of an extruder and by limiting the size of the polymer and porogen particle sizes used to fabricate the scaffold, final pore size can be controlled. This process also results in scaffolds with over 90% of pores being greater than 100 µm in size.

3.4.2 Proliferation

DNA content was measured on PCL scaffolds at each time point. The results of the assay performed on the seeded PCL scaffolds indicated that DNA content increased until day 14, and then decreased at day 21. The increase in proliferation was highest between day 7 and day 14 where a 165% increase in DNA was seen. A 29% decrease in cell number was observed between day 14 and day 21.These results are shown in Figure 3.8. This data is consistent with the previous studies using PCL scaffolds[32]

16000

14000

12000

10000

8000

6000 DNA DNA Content(ng) 4000

2000

Day 0 Day 7 Day 14 Day 21

Figure 3.8: DNA content of PCL scaffolds at various time points

3.4.3 SEM

SEM images of PCL/HA scaffolds and PCL scaffolds showed evidence of continued cell proliferation until day 14 and then a decrease was seen at day 21, similar to the results seen from the proliferation assay for the PCL scaffolds. Cells appeared to grow by forming flat, multi- layered structures on the scaffold. SEM images showed a much higher number of cells attaching to PCL/HA scaffolds compared to PCL scaffolds (Figure 3.9 and Figure 3.10). Previous studies have shown that marrow stromal cells and MC3T3-E1 cells adhere to bare PCL surfaces when cultured but do no support proliferation. [19, 33] The findings of this study would seem to support

this and suggest that HA increases the proliferation of cells seeded on the scaffolds. Cell bridging of scaffold pores can be observed at day 14 in PCL/HA scaffolds (Figure 3.9 d) suggesting that smaller pores could be covered by cells as proliferation continues over time.

Figure 3.9: SEM images of PCL/HA scaffolds: a) Blank Figure 3.10: SEM images of PCL scaffolds: a) Blank scaffolds with no cells b) day 0 c) day 7 d) day 14 e) day scaffolds with no cells b) day 0 c) day 7 d) day 14 e) day 21 21

3.5 Conclusions

The results of this study show that melt processing is a viable fabrication technique for creating structurally stable PCL and PCL/HA scaffolds. HA was successfully incorporated into the scaffold without negatively affecting scaffold morphology or pore formation. Pore size was controlled by limiting the size of the PCL and PEO particles used in the fabrication process. SEM images showed that the incorporation of HA into the scaffolds increased the cell growth and cell attachment seen over time on scaffolds seeded with MC3T3-E1 cells. Proliferation assays performed on PCL scaffolds showed an increase in DNA content up to day 14 suggesting that the scaffolds are capable of facilitating cell attachment and proliferation in vitro over time.

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15. Parsons, J.R. Cartilage. In: Black, J., and Hastings, G., eds. Handbook of Biomaterials Properties. New York: Chapman & Hall, 1998, pp. 40–46. 16. Woo, S.L.-Y., and Levine, R.E. Ligament, tendon and fascia. In: Black, J., and Hastings, G., eds. Handbook of Biomaterials Properties. New York: Chapman & Hall, 1998, pp. 59–65.

17. Yousefi AM, Gauvin C, Sun L, DiRaddo RW, Fernandes J. Design and Fabrication of 3D- Plotted Polymeric Scaffolds in Functional Tissue Engineering, Polymer Engineering and Science 2007;47:608-618.

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19. Washburn NR, Simon CG, Tona A, Elgendy HM, Karim A, Amis EJ. J Biomed Mater Res 2002;60:20–9.

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21. Maquet V, Jerome R. Design of macroporous biodegradable polymer scaffolds for cell transplantation. Mater Sci Forum 1997;250:15–42.

22. Peters MC, Mooney DJ. Synthetic extracellular matrices for cell transplantation. Mater Sci Forum 1997;250:43–52.

23. Ma PX, Zhang R. Synthetic nano-scale fibrous extracellular matrix. J Biomed Mater Res 1999;46:60–72.

24. Barbanti, S. H., Santos Jr., A. R., Zavaglia, C. A. C., & Duck, E. A. R. Poly(e-caprolactone) and poly(d,l-lactic acid-co-glycolic acid) scaffolds used in bone tissue engineering prepared by melt compression–particulate leaching method. J. Mater Sci: Mater Med 2011;22, 2377- 2385.

25. Shikinami, Y., and Okuno, M. Bioresorbable devices made of forged composites of hydroxyapatite (HA) particles and poly-L-lactide (PLLA): Part I. Basic characteristics. Biomaterials 20, 859, 1999.

26. Kim, J., Kim, I. S., & Cho, T. H, et al. Bone regeneration using hyaluronic acid-based hydrogel with bone morohogenic protein-2 and human mesenchymal stem cells. Biomaterials 2007;1830-1837.

27. Wagoner Johnson, A. J., & Herschler, B. A. A review of the mechanical behavior of cap and cap/polymer composites for applications in bone replacement and repair. Acta biomaterialia 2001;7, 16-30.

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Chapter 4

1. Significance

The goals of this project were to fabricate scaffolds suitable for bone tissue engineering, optimize these scaffolds with regards to mechanical properties and porosity, and to finally assess the potential effectiveness of these scaffolds using mechanical testing, imaging, and biological assays. Progress has been made into accomplishing these goals, and the results of this study represent important steps toward the use of these scaffolds for practical applications.

The results of this study showed the melt processing was an effective fabrication technique for the creation of scaffolds with controlled pore size. The use of PEO as a porogen allowed it to be leached from the final scaffold by simply using water, removing the need for any toxic solvents. Once the PEO was completed removed, stable scaffolds with interconnected pores were formed. The final size of the pores formed could be controlled by controlling the particle size of the powders used in fabrication. Additionally, scaffolds were fabricated with over 90% of pores being greater than 100 µm in size. Both polymer and polymer/ceramic scaffolds were created using this technique and it was shown that the addition of HA into the scaffold had no negative effects on mechanical properties or porosity.

This work also showed that the properties of fabricated scaffolds could be tailored based on several parameters including polymer ratio, HA concentration, salt content and particle size, and applied pressure. This would allow scaffolds to be made with properties to match the specific target tissue they would replace. This would also allow a design of experiment (DOE) to be performed to optimize certain properties based on the type of tissue and the specific need.

Biological assays were done to examine the effectiveness of these scaffolds at facilitating cell growth. Proliferation assay on the scaffolds seeded with osteoblasts showed an increase in DNA content up to day 14 with readings at day 21 still being higher than day 0. SEM images confirmed these findings and revealed an increase in cell attachment and proliferation in PCL/HA scaffolds compared to PCL scaffolds. Images also showed how cells grow on the scaffolds as well as their ability to bridge pores over time. All of these results together show that the scaffolds are capable of facilitating cell growth in vitro over time.

2. Future Work

While significant progress was made in analysis and optimization of these scaffolds, future work could build on the groundwork laid by this study. An average Young’s Modulus of approximately 2 MPa was achieved with the scaffolds in this study which is significantly less than the 20-500 MPa required to match cancellous bone. However, the possibility of using these scaffolds for other applications such as bone graft substitutes to fill defects that are not under direct load should be explored. A design of experiment could possibly be used to aide in this research. Future biological assays should use the finalized protocols used in this study to examine additional cell growth trials to determine the variability between trials of the same scaffolds. This was not explored in this study and would be useful when comparing trials for different formulations of scaffolds.