Applied Science, Faculty of
Mechanical Engineering, Department of
DSpace
UBCO
Samanipour, Roya
Doctor of Philosophy - PhD
University of British Columbia
Nervous system disorders including acute traumatic injuries, neurodegenerative diseases, and neurodevelopmental disorders are estimated to affect more than one billion people worldwide. Study and understanding the development of the human nervous system and exposing the mechanisms of mental disorders has greatly been limited due to the restricted access to the functional human brain tissue. 3D in vitro organ models has recently shown to be a powerful tool for biological and medical studies. These models, however, require special 3D construction of cell and extracellular matrixes that are often hard to achieve with conventional fabrication approaches. Bioprinting technique has emerged as potent platform to fabricate these complex 3D models. Here, state-of-art stem cell-based 3D in vitro brain models that recapitulate the geometrical complexity of the brain are developed using 3D bioprinting. The model is developed based on two cell types, neural stem cell and primary astrocytes. To create the model, a high-throughput biofabrication strategy based on embedded 3D bioprinting technology is designed, developed and characterized. Protocols, culture media, bioinks and biomaterials used are tuned and optimized to increase cell viability, enhance cell activity and promote neural stem cell differentiation. The procedures are optimized through a series of 2D and 3D studies and finally, the 3D bioprinted brain in vitro model is carried out. Keywords: Brain tissue, Neural stem cells, Bioprinting, Differentiation
i Development of 3D functional brain tissue model by Roya Samanipour B.Sc., The Shahid Chamran University, 2007 M.Sc., The Tarbiat Modares University, 2012 M.Sc., The University of British Columbia, 2015 A THESIS SUBMITTED IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY in THE COLLEGE OF GRADUATE STUDIES (Mechanical Engineering) THE UNIVERSITY OF BRITISH COLUMBIA (Okanagan) December 2019 © Roya Samanipour, 2019 ii The following individuals certify that they have read, and recommend to the College of Graduate Studies for acceptance, a thesis/dissertation entitled: Development of 3D functional brain tissue model submitted by Roya Samanipour in partial fulfillment of the requirements of the degree of DOCTOR OF PHILOSOPHY . Dr. Mina Hoorfar (Supervisor) Dr. SuRyon Shin (Co-supervisor) Supervisor Dr. Homayoun Najjaran Supervisory Committee Member Dr. Keekyoung Kim Supervisory Committee Member Dr. Richard Klukas University Examiner Dr. Khashayar khoshmanesh External Examiner iii Abstract Nervous system disorders including acute traumatic injuries, neurodegenerative diseases, and neurodevelopmental disorders are estimated to affect more than one billion people worldwide. Study and understanding the development of the human nervous system and exposing the mechanisms of mental disorders has greatly been limited due to the restricted access to the functional human brain tissue. 3D in vitro organ models has recently shown to be a powerful tool for biological and medical studies. These models, however, require special 3D construction of cell and extracellular matrixes that are often hard to achieve with conventional fabrication approaches. Bioprinting technique has emerged as potent platforms to fabricate these complex 3D models. Here, state-of-art stem cell-based 3D in vitro brain models that recapitulate the geometrical complexity of the brain are developed using 3D bioprinting. The model is developed based on two cell types, neural stem cell and primary astrocytes. To create the model, a high-throughput biofabrication strategy based on embedded 3D bioprinting technology is designed, developed and characterized. Protocols, culture media, bioinks and biomaterials used are tuned and optimized to increase cell viability, enhance cell activity and promote neural stem cell differentiation. The procedures are optimized through a series of 2D and 3D studies and finally the 3D bioprinted brain in vitro model are carried out. Keywords: Brain tissue, Neural stem cells, Bioprinting, Differentiation iv Lay Summary In this thesis a brain tissue model is developed using bioprinting technique. This model could benefit understanding of neurodevelopmental mechanism. This models consists of two type of cells. The cells are spatially organized in a 3D structure using a bioprinter platform. The material which is used for this study is hydrogel material which is biocompatible material. The cells are grow in this biocompatible hydrogel and proliferate and get mature. Different combination of material is used for each cell type since every cell is need different environment in terms of stiffness to grow and proliferate. v Preface The research presented in this thesis is the original work performed by the author. This thesis was supervised by Dr. Mina Hoorfar at the Advanced Thermo-Fluidic Laboratory (ATFL) in the school of engineering, University of British Columbia and Dr. Su Ryon Shin at the Shin Laboratory in the Harvard Medical School. Parts of this thesis have been published in peer reviewed scientific journals and conference proceedings, and my supervisor is the co-authors in all of the publications. Contributions The results of this thesis were published in peer-reviewed journals and conference proceedings. The details of the author contribution is explained for each publication below: a. Refereed journal publications a. Roya Samanipour, Moritz Werb, Ting Wang, Hamed Hssanzad, Juan Ma Ledesma Rangel, Mina Hoorfar , Anwarul Hasan, Chang Kee Lee, Su Ryon Shin, “Ferritin Nano-cage Conjugated Hybrid Hydrogel for Sustained Drug Delivery and Tissue Engineering Applications.” ACS Biomaterials Science & Engineering, 2019. (Lead author, generating the basic idea, experiment and fabrication of the device, writing the paper). Copyright has been gotten. b. Afsoon Fallahi, Serena Mandla,Thomas Kerr-Phillip, Jungmok Seo, Raquel O. Rodrigues, Yasamin A. Jodat, Roya Samanipour, Mohammad Asif Hussain, Chang Kee Lee, Hojae Bae, Ali Khademhosseini, Jadranka Travas-Sejdic, and Su Ryon Shin, “Flexible and Stretchable PEDOT-coated Hybrid Substrates for Bioengineering vi Applications.” Chem Nano Mat 5 (6), 729-737, 2019. (Co-author, contributed in the development of the basic idea and device fabrication, and participated in writing the paper). c. Roya Samanipour, Hojat Rezaeinejad, Minoru Hirano, Su Ryon Shin, Mina Hoorfar, “A review on 3D printing functional brain model.” Submitted to Biomaterials. (Lead author, generating the basic idea, experiment and fabrication of the device, writing the paper). d. Roya Samanipour, Yi-Chen Li, Yasamin A. Jodat, Giulio Zorzi, Kai Zhu, Minoru Hirano, Ting Zheng, Karen Chang, Adnan Arnaout, Mina Hoorfar, Manu S. Mannoor, Su Ryon Shin, Ali Khademhosseini, “A 3D Embedded Co-culture Bioprinted Brain-like structure with Astrocyte-Neuron interactions,” under submission. (Lead author, generating the basic idea, experiment and fabrication of the device, writing the paper). e. Pooria Mostafalu, Hojatollah Rezaeinenejad, Mohamadmahdi Samandari, Roya Samanipour, Batzaya Byambaa, Ommar Soliman, Huseyin Avci, Karen Abrinia, Sinha Indranil , Ali Khademhosseini, Ali Tamayol, “Core-shell hydrogel fibers with tunable electrical and mechanical properties for engineering muscle tissue.” under revision. (Co-author, contributed in the development of the basic idea and device fabrication, and participated in writing the paper). f. Mohamed G. A. Mohamed, Pranav Ambhorkar, Roya Samanipour, Yang, Annie Ali Ghafoor, Keekyoung kim, “Microfluidics-based Fabrication of Cell-laden Microgels”, submitted to Biomacromolecules. (Co-author, contributed in the vii development of the basic idea and device fabrication, and participated in writing the paper). b. Conference proceedings and presentations g. Roya Samanipour, Hojatollah Rezaeinejad, Su Ryon Shin, Mina Hoorfar, “Digital-Micromirror-Device printed scaffold for vascularized tissue” MicroTAS 2018, Kaohsiung, Taiwan, Nov. 2018. (Lead author, generating the basic idea, experiment and fabrication of the device, writing the paper). h. Roya Samanipour, Hojatollah Rezaeinejad, Su Ryon Shin, Mina Hoorfar, “Developing a neuronal network-like platform on chip for drug test.” BMES 2018, Philadelpia, USA, Oct.2019. (Lead author, generating the basic idea, experiment and fabrication of the device, writing the paper). viii Table of Content Abstract .............................................................................................................................. iii Lay Summary ..................................................................................................................... iv Preface................................................................................................................................. v Table of Content .............................................................................................................. viii Table of Figures ............................................................................................................... xiii Acknowledgment ........................................................................................................... xxiii Dedication ...................................................................................................................... xxiv Chapter 1 Background ........................................................................................................ 1 1.1. Introduction .......................................................................................................... 1 1.2. 3D neural model ................................................................................................... 2 1.3. 3D brain model fundamentals and design factors ................................................ 5 1.4. State-of-the-art in development of 3D brain tissues............................................. 8 1.5. Motivation .......................................................................................................... 13 1.6. Objectives ........................................................................................................... 14 1.7. Thesis organization ............................................................................................ 16 Chapter 2 Fabrication of 3D Brain Model ........................................................................ 18 2.1. Defining the brain model.................................................................................... 18 2.2. Model and methodology..................................................................................... 21 ix 2.3. Experimental procedure ..................................................................................... 23 2.4. Material Preparation ........................................................................................... 29 2.4.1. Supporting bath hydrogel .............................................................................29 2.4.2. Cell suspended supporting bath ...................................................................29 2.4.3. Bioinks .........................................................................................................29 2.4.4. Cell-Laden Bioinks ......................................................................................30 2.4.5. Synthesis of GelMA (Gelatin methacrylate) ................................................30 2.5. Protocols and procedures ................................................................................... 31 2.5.1. Cell cultures and neural stem cell differentiation ........................................31 2.5.2. Embedded bioprinting procedures ...............................................................31 2.5.3. Viability and metabolic activity assays........................................................32 2.5.4. F-actin/dapi staining.....................................................................................33 2.5.5. Immunochemistry ........................................................................................33 2.5.6. Statistical analysis ........................................................................................34 2.5.7. UV Crosslinking ..........................................................................................34 2.5.8. Mechanical testing .......................................................................................35 2.5.9. Rheological Characterization .......................................................................35 Chapter 3 Development of Printing Biomaterial .............................................................. 36 3.1. Development of 3D bioprinting material ........................................................... 36 3.2. Optimizing Supporting bath ............................................................................... 36 x 3.3. Optimizing bioink and 3D printing parameter ................................................... 47 Chapter 4 Cell Study of Developed Biomaterials ............................................................. 57 4.1. Cell study of bath and printed cell-laden construct ............................................ 57 4.2. Optimizing supporting bath for cellular activity of astrocyte ............................ 57 4.3. Bioink Optimization for cellular activity of neuron cells .................................. 66 4.4. Summary ............................................................................................................ 70 Chapter 5 Neural Stem Cell Differentiation ..................................................................... 71 5.1. Neural stem cell differentiation into neuron ...................................................... 71 5.2. Stem cell differentiation in 2D ........................................................................... 71 5.3. Differentiation study in 3D printed constructs ................................................... 74 5.4. Summary ............................................................................................................ 81 Chapter 6 Co-culture ......................................................................................................... 82 6.1. Co-culture of Neuron stem cell and Astrocytes ................................................. 82 6.2. Stage 1: optimization of the differentiating media for astrocyte growth in 2D . 83 6.3. Stage 2: optimization of the media in 2D co-culture ......................................... 83 6.4. Stage 3: optimization of the media in 3D co-culture ......................................... 86 6.5. 3D bioprinted brain model co-culture ................................................................ 87 6.6. Summary ............................................................................................................ 89 Chapter 7 Conclusion & suggestions for future work ...................................................... 90 7.1. Summary ............................................................................................................ 90 xi 7.2. Contribution to the field ..................................................................................... 90 7.3. Future work ........................................................................................................ 91 Appendix A Engineering biocompatible hydrogel fibers with tunable mechanical properties for neural tissue engineering ............................................................................ 93 A.1. Introduction ........................................................................................................ 93 A.2. Material and Methods......................................................................................... 95 A.3. Results and discussion ........................................................................................ 97 A.3.1. Fabrication of Hybrid Hydrogel Fiber .............................................................97 A.3.2. Physical properties ...........................................................................................99 A.3.3. Cytotoxicity evaluation ..................................................................................101 A.4. Conclusions ...................................................................................................... 104 Appendix B Ferritin Nano-cage Conjugated Hybrid Hydrogel for Sustained Drug Delivery and Tissue Engineering Applications .............................................................. 106 B.1. Introduction ...................................................................................................... 107 B.2. Material and methods ....................................................................................... 111 B.3. Results and Discussion ..................................................................................... 116 B.3.1. Synthesize and characterization of methacrylated ferritin and apoferritin 116 B.3.2. Morphological evaluation of nanoparticle composite GelMA hydrogels..117 B.3.3. Mechanical Properties of nanoparticle composite GelMA hydrogels .......120 B.3.4. Cell encapsulation in nanoparticle composite GelMA hydroels................128 xii B.3.5. FITC release from GeMA-Apoferritin and GelMA-ApoMA encapsulated FITC..........................................................................................................................132 B.4. Conclusion ........................................................................................................ 136 Appendix C G-Code for cellink bioprinter ................................................................ 139 Bibliography ................................................................................................................... 151 xiii Table of Figures Figure 1.1. The motivation behind development of the brain models. ............................... 2 Figure 1.2. Some of the current neural models and their main weaknesses ....................... 4 Figure 1.3. Complexity of the neural system and the basics to consider when designing a brain model. .................................................................................................. 8 Figure 3.2. Schematic process of the dynamic structure change of the GelMA solution under a shear tress. ........................................................................................ 40 Figure 3.3. The effect of calcium ions on the viscosity of the GelMA solution with the different concentrations. ................................................................................ 41 Figure 3.5. The degradation behaviors of 5 wt.% chemical GelMA supporting hydrogel after 7 and 14 days of incubation. .................................................. 43 Figure 3.6. a) Oscillatory strain sweep between low (1%) and high (250%) strain shows 5% physical GelMA hydrogel with a rapid self-healing ability during the printing process. b) The scratch recovery of 5% GelMA after the movement of the nozzle, showing that the scratch was recovered within 30 min. ................................................................................................ 45 Figure 3.7. (a) Schematic process showing the self-healing behavior of GelMA baths under shear stress triggered upon the needle movement. (b) SEM images of 5% GelMA bath before (left) and after (middle) the movement of nozzle in the bath showing the self-healing process (right) of the cracks in the bath .......................................................................................................... 46 Figure 3.8. Schematic diagram of the embedded bioprinted GelMA-based bioink ......... 47 xiv Figure 3.9. a) Continuous flow of 2% alginate mixed with 7%, 5%, 3.5%, and 3% gelatin tested as bioinks at an equivalent shear rate. The results show that the bioinks have the shear-thinning behavior. b) Oscillatory strain sweep between low (1%) and high (250%) strain shows the rapid phase transition from a gel-like to fluid-like behavior for the 5% gelatin/2% alginate bioink, i.e., necessary for localization of extruded bioinks. .......................... 49 Figure 3.10. Oscillatory stress model shows the storage modulus (G’) of bath-5% is higher than other inks conditions, it suggests that the 5% GelMA as a bath is with ability to maintain the fiber structure after printing. 3, 4, and 5% inks represent 3, 4 and 5%GelMA in the inks. All inks include 1% gelatin, and 2% alginate. ............................................................................................ 50 Figure 3.12. The morphologies of fibers were printed from the a) 27G and b) 30G nozzles at the various extrusion rates and fixed nozzle moving speeds. ....... 53 Figure 3.13. The circularity of printed fibers extruded from 30G nozzle in the supporting gel under various nozzle moving speeds and extrusion rates. ..... 54 Figure 3.14. Phase contrast imaging of the extruded parallel fibers with circular cross-section and approximately 300 μm thickness ................................................ 55 Figure 3.15. Young’s modulus of different GelMA ink concentrations. 3, 4, and 5% inks represent 3, 4 and 5%GelMA in the inks. All inks include 1% gelatin, and 2% alginate. The stiffness of our optimized condition (ink 4%) is in the range of brain stiffness. ........................................................................... 56 Figure 4.1. The morphologies of astrocytes in the 3.5%, 5%, and 7% chemical GelMA gel with different UV exposed times after 14 days of incubation. It is xv shown that the astrocytes could extend and have spindle morphologies in 5% chemical GelMA hydrogel under the 60s UV curing time. .................... 58 Figure 4.2. Live/dead viability assay of astrocytes in 5% GelMA bath upon different UV exposure treatments. ............................................................................... 59 Figure 4.3. GFAP/ DAPI staining of astrocytes in the 5wt% chemical GelMA supporting gel after a) 7 days of incubation, and b)14 days of incubation. .. 60 Figure 4.4. Schematic of the crosslinked supporting gel containing astrocytes: a) the structure without pillars, b) the structure with posts. .................................... 61 Figure 4.5. Live/dead assay for MCF-7 at day 7: a) live/dead assay for the structure with pillars, b) live/dead assay for the structure without pillars. The images were taken by a florescent microscope. The blur part shows that the cells are in different planes of focus. ..................................................................... 62 Figure 4.6. Actin/ DAPI staining of 3T3 cells at different days: a) day1, b) day2, c) day3. The images were taken by a florescent microscope. The blur part shows that the cells are in different plane focus. ........................................... 63 Figure 4.7. a) The live/dead assay and cell viability of astrocytes in the 5% chemical GelMA supporting gel after UV treatment. b) The cell viability percentage of astrocytes at different UV exposure times. c) The metabolic activity of astrocytes for different UV exposure times at three different time points (1 day, 7 days, and 14 day). ........................................................................... 65 Figure 4.8. GFAP/ DAPI staining of astrocytes in the 5wt% chemical GelMA supporting gel after a) 7 days of incubation, and b) 14 days of incubation. ....................................................................................................................... 66 xvi Figure 4.9. Actin/ DAPI staining of 3T3 cells encapsulated in the GelMA/alginate ink and treated at a) 70 s, and b) 80 s UV after 14 days of incubation. ............... 68 Figure 4.10. (a) Phase contrast images and (b)confocal images of live/dead staining of neural stem cells in the printed fibers after embedded printing in the 5% GelMA bath. .................................................................................................. 69 Figure 4.11. Phase contrast images of (a) morphology and (b) F-actin/DAPI staining of neuron cells in the printed fibers after embedded printing in the 5% GelMA bath. Confluent parallel fibers with a thickness of around 300 μm are created upon 7 days of in vitro culture. ................................................... 70 Figure 5.1. Optimizing the differentiation protocol for neuronal differentiation in (a) 2D and (b) 3D. RA in defined medium (MEM/F12, 1% ITS, P/S) was added for 48 hours. At this stage, neural stem cells dominate the culture in dense planar sheets of culture (stage 1). After RA removal, dense aggregated colonies start to form in the culture (stage 2, a(i), b(i)). In 2D, these aggregates are loose and gradually detach from the surface, forming floating neural spheroids. In 3D, the spheroids form inside the printed fibers through migration and aggregation of cells (stage 3, a(ii), b(ii)). After 7-8 days of in vitro culture, the aggregated spheroids settle in the 2D culture and attach to the surface followed by neurite formation (stage 4, a(iii), b(iii)). Starting day 10 of differentiation, the neurites and neural networks are formed in the 2D culture. In the 3D culture, the neurites and neural networks are formed across and around the neuro-spheroids. Neurons are further matured by axonal development and high expression xvii of specific biomarkers such as beta III tubulin (stage 5, a(iv), b(iv)). If FBS-containing culture media is supplied to the culture, the neurites fasciculate and the population will gradually get populated by neuroglia (stage 6, a(v), b(v)). To avoid this condition, neuron maturation media such as neurobasal medium or defined medium can be used. Scale bar in all images is 200 μm. ..................................................................................... 73 Figure 5.2. Beta tubulin/DAPI staining of neuron cells after 7 days of culture. a) DAPI channel, b) Beta tubulin channel, c) merged channel. ................................... 75 Figure 5.3. a) F-Actin/DAPI staining of the neuron cells after 14 days of culture. b) Beta tubulin staining of the neuron cells after 14 days of culture. ................ 75 Figure 5.4 Metabolic activity of neural cell-laden bioink over the 14-day culture period ....................................................................................................................... 77 Figure 5.5. Fluorescent images of neural cell-laden fibers labeled by F-actin and DAPI after day 7. ..................................................................................................... 77 Figure 5.6. Studying the effect of Alginate concentration in the bioink on neuron specific marker expression. Three ink conditions of alginate were printed and imaged after 10 days of differentiation. So significant difference was observed between the three. Alginate 0.5% was chosen for its best printability. Inks containing Alginate with a concentration above 1% showed compromised printability and were not selected. ............................. 78 Figure 5.8. (a) phase contrast image of the neuron-laden fiber 20 days after neural differentiation induction, showing neuron spreading (b) βIII tubulin and DAPI staining of the aligned printed fibers (boundaries shown in red) after xviii 14 days of culture on day 10 of differentiation. (c,d) an internal network of neurons in the printed nerve fibers expressing positive βIII tubulin and negative GFAP which shows successful differentiation of neural stem cells into neurons and not neuroglia. ..............................................................81 Figure 6.1. Optimization of co-culture media for brain-like co-culture constructs. Astrocytes were cultured in three different media known to enhance neuronal differentiation. Neurobasal medium supplemented with B27 (NBM), Defined Medium (DM, containing MEM/F12+1%ITS+1%P/S), and neural stem cell media (NS, containing MEM+10% FBS+1% P/S) were tested. DMEM astrocyte media (DMEM, 10% FBS, 1% P/S) was selected as the control culture. Astrocytes cultured in NBM did not survive the one-week culture. Cells cultured with DM, however, showed good proliferation. As such, DM was selected as the common media for co culture of neuron- and astrocyte- laden constructs. GFAP/DAPI staining (right column) confirmed this observation. Scale bar in all images is 200 μm. ................................................................................................................. 85 Figure 6.2. Co-culture fluorescent images of neuron and astrocyte cells labeled by BIII tubulin and GFAP individually on 2D after 14 days of incubation ............... 86 Figure 6.3 Co-culture fluorescent images of neuron and astrocyte cells labeled by BIII tubulin and GFAP individually in a thin layer of hydrogel 14 days of incubation. ..................................................................................................... 87 xix Figure 6.4. Co-culture fluorescent images of neuron and astrocyte cells labeled by BIII tubulin and GFAP individually in the brain-like structures after 14 days of incubation ...................................................................................................... 88 Figure A.1. Conceptual view of the high throughput fabrication process of core-shell hydrogel fiber and the resulted construct. (A) The setup was consisted of a microfluidic device with the ability to control the flow rate of core and shell solution, separately. Extruding the solutions through co-axial microfluidic device can induce shear force within the fluid and elongate polymer (i and ii). The structure then was injected to CaCl2 bath where alginate was crosslinked to form the fiber’s matrix. Crosslinking alginate trapped the hydrogel in the core (GelMA/gelatin) and formed a template network in the shell (iii). (B) Finally, the samples were exposed to UV-irradiation for crosslinking GelMA. Representative micrographs showing produced fiber. Scale bar is 500μm. (C) schematic of neural fiber generation, alginate acts as a template preventing neurons from diagonal expansion and force them to within the GelMA/gelatin core and along the fiber, which leads to interconnection between neurons. ................................ 99 Figure A.2. Characterization of hydrogel fibers. (A) fabrication setup of core-shell hydrogel fibers (B) Geometrical features of core-shell hydrogel fibers as compared to sold hydrogel fiber (C) The diameter can easily be controlled by setting the proper inflow rate of sheath and core fluids (D) Mechanical properties different alginate concentration (E) Mechanical properties of xx GelMA/gelatin hydrogel with variation of GelMA concentration, gelatin was kept constant as 1% for all (*: P<0.05) ................................................ 101 Figure A.3. Viability and proliferation of neuroblastoma encapsulated in fiber. (A) Phase-contrast images of neuroblastoma encapsulated in hydrogel fiber in day 0, day 1, and day 5, scale bar 100 um. (B) Live/ dead staining of neuroblastoma at day 1 and day 3, scale bar 100 um. (C) F-actin/DAPI staining at different days of culture, day 1, day 5, day7, and day14, scale bar 50 um. .................................................................................................... 102 Figure A.4 Qualitative and quantitative analysis of cellular orientation in different days of the culture for the encapsulated cells in fiber. (A) F-actin/DAPI staining after 7 and 14 days, respectively from left to right. (B) Deviation of neuroblastoma from the channel direction for Days 7 and 14. (C) Cell seeding density and culture density after Days 7 and 14............................. 104 Figure B.1. Characterization of methacrylated ferrtin/apoferritin. (a) Schematic of Ferritin methacrylation (b) Schematic of Ferritin methacrylation (c) Methacrylation degree of ferritin and apoferritin. (d) Transmission Electron Microscope (TEM) of (i) Ferritin, (ii) FerMA, (iii) Apoferritin, and (iv) ApoMA. Statistical analyzes have been performed, * p < 0.05. .... 117 Figure B.2. Synthesis and characterization of ferritin/apoferritin-GelMA hydrogels. (a) Schematic of Ferritin methacrylation with GelMA (b) Optical images of ferritin-GelMA and apoferritin-GelMA prepolymer with different concentration of ferritin and apoferritin (c) UV-Vvis absorption spectra of ferritin-GelMA and apoferritin-GelMA prepolymer solutions with xxi different concentrations of ferritin and apoferritin, grey range (320 nm-390 nm) indicate the wavelength of UV curing system. (d) Scanning electron microscopy (SEM) of 5 wt% GelMA only, 5 wt% GelMA- 0.1 mg/ml ferritin, and 5 wt% GelMA- 0.1 mg/ml FerMA. Scale bar is 150 nm. ............................................................................................................... 119 Figure B.4. Viability and elongation of cells encapsulated in nanoparticle incorporated composite hydrogels. (a) Live dead assay images (b) Quantification live/dead (c) Phase contrast images 3T3 cell encapsulated in GelMA -ferritin and GelMA- FerMA, for 10 sec UV exposure time at day 2 of culture. (d) Quantification chart of speaded cells in GelMA -ferritin and GelMA- FerMA for 10 sec and 20 sec UV exposure time at day 2. (e) Actin/DAPI staining of 3T3 fibroblast encapsulated in 5 wt% GelMA hydrogels with different concentration of ferritin and FerMAafter 7 days of culturing (UV exposure time 10s). DAPI (blue) for nuclei, phalladoin (green) for f-actin. Images taken with 100x magnification, white scale bar equals to 200 μm, (f) Percentages of actin covered with filamentous actin. (g) Actin/DAPI staining of 3T3 fibroblast encapsulated in 5 wt% GelMA hydrogels with different concentration of apoferritin and ApoMA after 7 days of culturing (UV exposure time 10s). DAPI (blue) for nuclei, phalladoin (green) for f-actin. Images taken with 100x magnification, white scale bar equals to 200 μm, (h) Percentages of actin covered with filamentous actin. Statistical analyzes have been performed, (data set that xxii show not significantly difference (ns. ;> 0.05), data set that are at significant levels: * p < 0.05, ** p < 0.01, *** p < 0.001. .......................... 130 Figure B.5. Drug delivery potential of apoferritin/ApoMA-GelMA hydrogels. (a) Schematic of capturing FITC with Apoferritin. (b) Absorbance spectra of plain FITC (red color), Filtered FITC captured in Apoferritin (blue color), Filtered FITC mixed with Apoferritin (green color), and Apoferritin (black color) were measured. (c) Florescence images of (i) Plain FITC, (ii) Filtered FITC captured in Apoferritin, and (iii) Filtered FITC mixed with Apoferritin. Scale bar is 2mm. (d) FITC release profile of FITC laden GelMA-Apoferitin and GelMA-ApoMA. ................................................... 135 Figure B.6. UV-Vvis absorption spectra of FerMA-GelMA and ApoMA-GelMA prepolymer solutions with different concentrations of FerMA and ApoMA, grey range (320 nm-390 nm) indicate the wavelength of UV curing system. .............................................................................................. 137 Figure B.7. Actin/DAPI staining of 3T3 fibroblast encapsulated in 5 wt% GelMA hydrogels with 10 sec UV exposure time after 7 days of culture. DAPI (blue) for nuclei, phalladoin (green) for f-actin. Images were taken with 100x magnification, ..................................................................................... 137 Figure B.8. Phase contrast images of 3T3 fibroblast encapsulated in ferritin-GelMA and FerMA-GelMA with 10 sec UV exposure time after 2 days of culture. ..................................................................................................................... 138 xxiii Acknowledgment In the first place, I would like to express my gratitude to my supervisor Dr. Mina Hoorfar for her great guidance and supervision during my PhD career. She has always been a great role model and source of inspiration to me. I would like to thank my co-supervisor, Dr. Su Ryon Shin for her support and guidance. I would like to thank my committee members Dr. Najjaran and Dr. kim for their kind advisory throughout my research. I am also grateful to those who helped me with their greatest knowledge and provided me with their resources. I am very thankful to all of my friends and colleagues who shared part of their lives with me and created unforgettable memories. I would like to thank my family for their continued love and support. I am forever grateful for their generous time and compassion. Lastly, I would like to thank my husband, Hojatollah Rezaeinejad, for being there every step along the way. The journey would not have been the same without him. Words cannot describe my appreciation for his endless insight. xxiv Dedication This thesis is dedicated to my family, who have dedicated their true love and support to me. There is no word to express my deepest gratitude and love to them and I and very grateful to have such wonderful people in my life. 1 Chapter 1 Background 1.1. Introduction Disorders in the nervous system are estimated to affect more than one billion people worldwide. The neural system has complex 3D architecture in the form of fibers with high level of connectivity and contains various cell types. The central nerve system (CNS) itself contains approximately 86 billion neurons and 85 billion non-neural [1], known as glial cells. Glial cells can be categorized into astrocytes, microglia, oligodendrocyte, ependymal cells, and microglia [2]. Progress in understanding the human nervous system and elucidating the mechanisms of neurodevelopmental disorders has greatly been limited due to the restricted access to the functional human brain tissue. Significant amount of effort are therefore focused to develop a valid in vitro 3D neural model that can simulate to some extend the complexity of the native tissue and could serve as a mean to study neural disorders, to evaluate effectiveness of the new medications, and to fundamentally understand and simulate neural functions and biology (Figure 1.1). More recently, emerge, development and utilization of 3D printing technology to produce 3D live tissue constructs (more known as bioprinting technology), and development of variety of printable biomaterials have facilitated biofabrication of complex tissue constructs in which were hard to achieve before[3-9]. Recently, bioprinting technology has also been utilized to fabricate 3D brain complex tissue models. In the following, 3D in vitro neural models are discussed, the fundamental design factors to create a model are explained and the state-of-art of neural model are reviewed. Next, the motivation and objectives of this thesis are 2 described, and finally, a novel brain model and biofabrication strategy based on 3D bioprinting to create the model are proposed. Figure 1.1. The motivation behind development of the brain models. 1.2. 3D neural model 3D in vitro neural models have recently received significant attention in the field of the modern neuroscience for studying neural circuitry, nerve regeneration, and neural disease. Functional and effective 3D neural engineered tissue models, in fact, can provide better understanding of brain development, facilitate exploration of new therapeutic options for the central nerve system (CNS) disorder, and enable cost-effective drug 3 discovery investigation and toxicology evaluation. In long term, 3D in vitro neural models have the potential to represent the human neural system and can be used in regenerative medicine to repair or replace a damaged part of human neural tissues. 3D in vitro models (in terms of strategy) can be categorized into cell-based models (e.g., spheroids, organoids), and engineered models (e.g., scaffold-based, microfluidics [10-16], and biotextiles). Spheroids and organoids can both be considered as cell-based models. Spheroids can be seen as non-adherent 3D cell cultures where their degree of heterogeneity is related to the initial population of cells. On the other hand, the organoids model has a higher order of assembly, and due to their organ-like structure, a matrix is required for their formation (e.g., matrigel). Cell-based models are superior at mimicking early developmental details. However, the engineered models can control the composition of the material and cell organization (for achieving tissue-like constructs), and therefore have more controlled and consistent outcomes. Regardless of the approach (cell-based or engineering approach) in the development of a miniaturized organ or a functional part of an organ, a user-controlled and replicable approach is essential for recapitulating the native neural tissue structure/organization and the microenvironment (Figure 1.2). 3D bioprinting has emerged to address the lack of flexibility in spatially positioning of the cells and to pattern cells into the desired complex architecture. 3D printing is an atomized high-throughput platform with exceptional versatility providing relatively fast fabrication of tissue constructs with complex 3D topologies. Different modes of 3D bioprinting (offering different capabilities and limitations) include inkjet, extrusion, and stereolithography printing. The main challenge in adapting printing technologies for biofabrication of a neural tissue model is the development of the bioink. To create a bioink for a specific tissue 4 type, a combination of biomaterial with printability and gelation property and materials with suitable biological properties (e.g., cell attachment, cell growth) compatible with the target tissue is conventionally used. Figure 1.2. Some of the current neural models and their main weaknesses There are an extensive range of biomaterials (mainly alginate, Poly (ethylene glycol) acrylate (PEGDA), fibrin, chitosan, hyaluronic acid, silk fibrin, gelatin, agarose, methylcellulose, and collagen) that has been used for bioprinting. Biomaterials must have two major characteristics to be considered for printing. These two characteristics are printability and gelation. Gelation might not be compatible with all printing technology, and since each of the mentioned biomaterials requires different conditions for the gelation process, they can only be used with their compatible printing setup to create a 3D printed construct. 5 As for the materials with suitable biological properties, hydrogels have been widely used to create in vitro neural constructs. Their tissue-like mechanical/chemical structure (such as their high porosity and low stiffness) and their physical/chemical properties (such high water content and their ability to be functionalized with bioactive components) makes this class of materials suitable for fabricating neural tissue models. Some of the most prominent hydrogels used for neural tissue culture include agarose[17-20], collagen type1[21-25], hyaluronic acid [26-28] PEG (Poly ethylene glycol)[27-29], chitosan [27, 30, 31], alginate [32-34], silk fibroin [35-39], and methylcellulose [40, 41]. Most of these hydrogels have basic characteristics of the neural cell culture including low stiffness, sufficient oxygen nutrition, waste diffusion, and cell attachment sites that do not irregulate neural phenotype [2]. 1.3. 3D brain model fundamentals and design factors Brain has a 3D structure where cells are positioned in a 3D extracellular matrix (ECM) with a set of defined physical properties and chemical compositions, where these physiochemical properties play a crucial role in neural cell-cell and cell-environment communications. Inevitably, the presence of the 3rd-dimension is essential for such a complex composition to exist, and simplifying such a composition into 2D models may cause aberration of neural phenotype [42]. However, even by considering those properties, there are still several biological/biochemical/biophysical elements (some known and some other still unknown) in both cellular and environmental levels that should be taken into account for building an organ. Hence, fabricating an entire organ is not practical in most cases. A more practical approach is to generate a 3D in vitro tissue model that may have 6 essential parts of the organ or a functional unit of an organ with sufficient complexity for representing the native tissue structure [2]. These 3D in vitro tissues should be evaluated based on their ability in recapitulating the bio-functionality of the native tissue. The 3D tissue models must also consider the physiological complexity of the native tissue in order to represent an accurate of model [43]. Here major fundamental features of the neural tissue was described as a guideline for designing neural tissue construct (Figure 1.3). One other major factor required for designing a neural construct is the composition of ECM used in the model. A neural system, in fact, has a unique ECM where the construct is mainly made of proteoglycans of lecticans, hyaluronan, and tenascins [44]. Interestingly, a component such as a collagen, fibronectin, and laminin are in much lower quantity if compared to another part of the human body [2]. ECM of the neural system also includes abundant soluble factors including growth factors, cytokines and chemokines, where these factors function in a concentration manner [45]. Neural development and circuit formation (including dendrite outgrowth and axonal targeting) function on finely tuned growth factor concentrations. Consequently, a key component to be included in vitro 3D models is to establish such concentration distribution to facilitate nerve regeneration [46-49]. Aside from the neural composition, neural ECM also may regulate cell behaviors through its physical and mechanical properties. Neural ECM has distinctive biophysical properties: it is known for its low elastic modulus and large porosity (compared to other tissues (such as heart, cartilage, and bones)). In particular, the elastic modulus of the brain tissue for neonatal is approximately 110 Pa, and for adults is less than 1 kPa [42]. It is also known that matrix stiffness has a significant effect on neural cell behaviors and morphologies [20, 29, 36]. For example, neural stem cells are more likely to differentiate 7 into glia when their surrounding matrix has Young’s modulus higher than 1 kPa. Softer ECM (100-500 pa), however, tends to promote cell migration and differentiation into neurons [50-53]. Studies have shown that mesenchymal stem cells express undergo neural differentiation (express neural gene) on soft substrates (1 kPa) where on stiffer matrices (10 kPa) tend to differentiate into glial lineage [54]. Similarly, ECM porosity can greatly affect cell behaviors and metabolism [54]. Large pore sizes (more than 1000 micrometer) facilitate appropriate nutrient exchange, and ECM with porosity similar to the native tissue results in better cell migration [55]. Neural cells have also shown to respond to microstructure and geometrical cues [56-60]. Such cues have proven to improve cell viability [61], migration [62], proliferation [50], and differentiation [63]. Topological cues have also shown to facilitate neurite outgrowth [33, 36, 64, 65]. The identification and recapitulation of these essential parameters (such as porosity, mechanical property, biocompatibility, biochemical and physical gradients, and multiple cell types) are helpful in reconstructing neural tissues. After determining the key factors for designing a 3D neuron model, a right technique needs to be identified for fabricating the model. Different techniques have been used to fabricate 3D neural models including hanging drop for fabricating self-biology models (spheroid and organoid) [2] and microfluidics [2]. These techniques have significantly enhanced our understanding of neural tissue development and neuropathology [2, 66-70]. However, they cannot be considered as high throughput techniques which are required for using the model for drug testing [2]. Moreover, the lack of precise spatial organization of cells, which is very crucial for signaling the physiologically relevant cues and can subsequently lead to generation of functional tissue model, is still a challenge. 8 More recently, bioprinting techniques have been developed and used to fabricate heterogeneous tissue model with great consistency [71]. In such techniques, cells are mixed with chemical cues and biomaterial, and structures are printed in a 3D environment. Therefore, the cells could be organized spatially in an engineered construct [72]. Figure 1.3. Complexity of the neural system and the basics to consider when designing a brain model. 1.4. State-of-the-art in development of 3D brain tissues To bridge the gap between the in vitro models and the actual human brain physiology, several studies have attempted to design 3D models by taking advantage of engineering techniques such as microfabrication [73], 3D patterned grooves[74], and microfluidics [75]. An example of this attempt is the study reported by Shin et al. [76] who developed an embryonic stem cell-derived neuron-based compartmental microfluidic 9 device. This device provided an arrangement for neuronal axons to traverse the microchannel. Moreover, they concluded that their device can be used for the comparison study of dissociated neurons in different test conditions. Park et al. [77] also fabricated a neurospheroid-based microfluidic chip with a constant fluid flow to mimic the fluids in the in-vivo interstitial space of the brain microenvironment. Using the developed neurospheroid-based chip, they have consequently shown that the effects of the interstitial level of slow and diffusion-dominant flow on removal of amyloid-β, i.e., a major factor contributing to Alzheimer’s disease, can be evaluated. In addition to the microfluidic systems, another technique (for the development of 3D brain-like models) has been performed by fabricating neurospheroid-like tissues through aggregation of neural cells or neuronal cell suspension (using techniques such as hanging droplets). Tanaka et al. [78] reported that human organotypic cultures may yield the closest representation of in-vivo human brain tissues outside the live patient body. Nevertheless, preparing human brain models for brain disorder studies are extremely sporadic and expensive due to the limited specimen availability and the human brain complexity [79]. As such, Lancaster et al. [80] and Pasca et al. [81] developed human pluripotent stem cell-derived 3D cerebral micro-organoids capable of recapitulating the human cortical development, required to obtain human brain-like tissues. Accordingly, a 3D biomimetic human brain model containing both neurons and neuroglia was indicated as a strong requisite to allow for precise replication of the cell-cell interactions in the brain, namely, astrocyte-neuron interaction. Bioprinting of the neural tissue has recently been developed. For instance, a two-step printing technique has been used to fabricate neural tissues. In such a technique [82-84], scaffolds are first printed, and then a layer of the cells is seeded on them. However, 10 with the recent advancements in bioprinting, cells mixed with the bioink have been printed to create more complex structures (such as the neural tissues) with spatial organization of cells in the construct. Unlike two-step printing, material and cells can be printed simultaneously to create heterogeneous tissue models. These models greatly benefit downstream high-throughput drug testing and clinical applications. A few prominent studies in this field are reviewed below: Qi Gu et al. [85] Fabricated a 3D neural tissue using human neural stem cells (hNSC). They used an extrusion printing technique in which polyacrylate-based bioink was mixed with hNSCs. The printing was used to create a rather common and simple reticular construct. Their bioink was comprised of alginate, carboxymethyl-chitosan (CMC) and agarose in which alginate and agarose were used to create the structural support. The agarose was mainly used to regulate bioink viscosity during printing and prior to gelation; whereas alginate was used to enable gelation after printing by immersing the printed construct in cation bath (e.g., calcium chloride solution). CMC (a water soluble derivative of chitosan) was used in their ink to conduct cell survival. They have shown hNSCs have good viability and continue to renew and proliferate in their printed construct for 10 days. After the initial 10 days of culture, they differentiated hNSCs in the construct for two weeks into glial cells and neuron cells. They also studied the pore size of the bioink (alginate(Al), carboxymethyl-chaisson (CMC), agarose (Ag)) used for the mini tissue in their work. They stated that different concentrations of carboxymethyl-chaitason (CMC) indicated variable porosity, with [85] 5% and 3.5% w/v CMC associated with a highly and sparsely porous surface respectively, and 2% or less w/v CMC gels associated with negligible to no pores. They also stated that the compressive Young's modulus of the optimized ink (5% weight 11 per volume (w/v) Al, 5% w/v CMC and 1.5% w/v Ag ) is 7.5 kPa. Their seeding density for 3D printing was 5M/ml and the maximum printing resolution that they could achieve with their optimized ink was 500m, which can be considered a low printing resolution. They used hNSC and performed in-situ differentiation to neuron and neuroglia after 10 days post printing. Tuj1 marker was used to label the neuron cells. GABA and GAD (glutamic acid decarboxylase) markers were used to label GABAergic neurons. OLIGO2 and GFAP were used to stain neuroglia cells [85]. The highlight of their work is differentiation of the hNSC into glia and neuron cells. However, the results of staining, showing that the hNSC are differentiated into two different cell types, are not promising. Also, the resolution of their printed construct is not high. Moreover, cells were printed with the hydrogel randomly which is another drawback of the model. Rodrigo lozano et al. [86] have demonstrated a multi-layered neural model, where it consists of discrete layers of primary cortical neural cells encapsulated in a hydrogel with cell viability of 75%. They proposed a neural ink for 3D printing the brain tissue comprised of only 0.5% gellan gum modified with RGD. The 3D printed model was achieved using a hand-held device with a coaxial nozzle configuration where the ink and the crosslinking solution were injected through the inner and outer nozzle, respectively. The gel was crosslinked once exposed to a calcium chloride solution. The printing cell density was 1million/ml. The novelty of their work is the development of the modified hydrogel along which the cortical neuron cells grow and elongate. However, the resolution of their fabrication method (i.e., the co-axial handheld device) is very low, and the resulting system is not high-throughput. Another drawback of their work is the use of one type of cells for their 3D model. 12 Daeha Joung et al. [87] have printed a 3D scaffold by sequentially depositing the scaffold ink and multiple cell-laden bioinks in a layer-by-layer manner to create multiple channels. They printed the 3D silicone scaffold as a base for the channels, then they printed the cell-laden matrigel (in a layer-by-layer manner) to create multiple channels for mimicking the parallel fibers like in the spinal cord. They printed two neural lineage stem cells (iPSC-derived spinal neuronal progenitor cells (sNPCs) and oligodendrocyteprogenitor cells (OPCs)) with an alternative dot arrangement. They tested a few materials (GelMA, Gelatin/fibrin, and matrigel) and showed that GelMA and Gelatin/fibrin are suitable for fibroblast but not for neurons. However, matrigel was shown to be a good choice for neurons. They also presented an outgrowth of neurons in their printed structure and performed calcium imaging. They directly printed sNPCs and OPCs onto the scaffold and cultured for 7 days. The outgrowth of axons with the presence of associated OPCs was detected within the printed microchannels using the NeuN nerun mature marker, and axon detected with the beta3III-tubulin antibody. These results showed that 3D bioprinted sNPCs have differentiated into neurons with extended axons propagating in a designed printed channel. Finally, the activity of these neuronal networks was tested and confirmed by physiological spontaneous calcium flux studies [87]. The novelty of their work is printing two neural lineage stem cells with alternative dot arrangement (for modeling the spinal cord). However, this method of printing can have leakage since the printed silicone channels are not solid. 13 1.5. Motivation 3D tissue models are essential for studying the nerve regenerative and neural tissue. 3D bioprinting has already revolutionized many fields of research and shown to have a significant potential for fabricating tissue models. Microfluidics is another technique that could be used in the assist of 3D bioprinting to fabricate 3D tissue models. However, very limited studies have been performed to incorporate 3D printing technology for biofabrication of the brain models. Most cases is still based on 2D brain models in which the brain structure is too simplified. This models lack studying the effect of cell-environment interaction. For a few 3D brain models developed so far merely a single type of neuron cells is used. Such models have the limitation of studying the effect of cell-cell interaction. A 3D brain model that considers brain complexity in cellular, structural, and geometrical levels is still not available. The aim of this project is to devise a sufficiently complex 3D brain model using a high-throughput bio-fabrication strategy that includes neuron and glial cells (non- neural cells) where neurons are positioned in a fiber-like structure very similar to the brain native tissue. For this purpose, I will develop and validate the brain model with such complexity and incorporate embedded biofabrication (a high-throughput technique) to fabricate the free form structure of the brain model containing two different type of cells (astrocytes and neuron cells). I also developed a microfluidic techniques to fabricate aligned neural fiber. This platform could be combined with bioprinting techniques (the co-flow nuzzle of bioprinting could be used as a nuzzle of bioprinting) to fabricate more complex 3D structures. 14 1.6. Objectives With the use of high-throughput embedded 3D bioprinting, this project aims at the development of a 3D tissue construct. With the use of neural and non-neural cells, a soft brain-like co-culture construct will be created by merely using polysaccharide biomaterials as the bioink and supporting gel. As such, a free-standing neuron-laden structure is directly fabricated in an astrocyte-laden supporting gel. The engineered brain-like biological construct is proposed to provide a cyto-compatible platform to recapitulate neuronal functions as well as opening pathways towards studying the complex neuron-astrocyte interactions in the brain. To create a construct that fairly represents the brain native tissue structure, several factors will be considered including the geometrical, cellular, and mechanical characteristics of the brain. The specific targets of this proposal are outlined as bellow: Aim 1 – To devise and develop a 3D bio-printing strategy for bio-fabricating a 3D construct that mimics a fiber-like structure of the brain. To realize this aim the following tasks are defined: a) Development and bio-fabrication of a brain model using embedded printing. b) Characterization and optimization of the bio-printing materials and printing ink for printability, crosslinkability (gelation), mechanical property and viscosity. Aim 2 – To develop and bio-fabricate a cell-laden construct based on the developed 3D printing technology (proposed in Aim 1) to incorporate the use of neural stem cells and astrocytes (non-neural cells). To realize this aim the following tasks are defined: 15 a) Characterization and optimization of the biomaterials biocompatibility, cell activity and proliferation, and 3D morphology of cells in hydrogel for both neuron and astrocytes. b) Optimization of the growth media for co-culture of the neural stem cells and astrocytes in 2D model. c) 3D printing the construct with biocompatible biomaterial developed, optimized and characterized in the previous step. Aim 3 – To develop and optimize a protocol for maturating the model in long-term culture to differentiate neural stem cells to neurons in the fabricated 3D model. To realize this aim the following tasks are defined: a) Optimization of the growth media and growth factors for differentiation of the neural stem cells into neurons in 2D. b) Optimization of biomaterial composition and mechanical properties for differentiation of the neural stem cells into neurons. c) Development of a 3D construct with neural stem cells and astrocytes co-cultured together, and differentiation of the stem cells into neurons. d) Biological characterization of the construct by studying factors such as cell viability, morphology, proliferation and migration. Aim 4 – To validate and study biological brain tissue representation of the fabricated and matured 3D brain model through immunostaining of Beta-tubulin and GFAP bio-markers. To realize this aim the following tasks are defined: 16 a) Immunostaining of the neurons in the 3D brain construct with Beta-tubulin to validate the differentiated neurons in the model and reveling the organization neurons in 3D. b) Immunostaining of the astrocytes in the 3D brain construct with GFAP to validate the presence of astrocytes in the model and their spatial distribution in 3D. 1.7. Thesis organization The organization of the thesis is as follow. The thesis consist of 6 chapters and 3 appendixes. The 1st chapter covers an introduction to neural 3D in vitro models and the motivation behind developing such models and motivation and objectives of this thesis. The chapter also compares different biofabrication strategies used to create in vitro brain models and embolden the use of bioprinting approach to create complex 3D tissue constructs. The 2nd Chapter of the thesis describes the details and novelties of the proposed brain model and explains the high throughput biofabrication strategy designed, here, to create the brain model. The chapter also covers several biomaterial preparation protocols and explains the procedures followed to characterize and study the model and the biomaterials. The 3rd chapter focuses on developing and physical characterization of the biomaterials required to carry out the biofabrication plan proposed in the 2nd chapter. Here, a series of physical characterizations were performed to optimize the printability of the 17 bioink and enhance the robustness of the supporting bath required for embedded printing strategy depicted in chapter 2. The 4th chapter is dedicated to cell study of the developed biomaterials in the chapter 3. The main purpose of this chapter is to validate the biocompatibility of the materials as well suitability of them for the specific cell type used in the model. Here, printing parameters were also optimized to increase cell viability and enhance cell activity both in the supporting bath hydrogel (for astrocytes) and within the printing bioink (for neural stem cells). The 5th chapter is focused on differentiating the neural stem cells toward neurons. The chapter describes the optimization of the cell media, tuning of biomaterial composition and possibly adding other biomaterials, and validating the results through immunostaining staining of differentiated neurons in the model. The 6th chapter is dedicated to present the co-culture results that includes both neural stem cells and astrocyte. Here the media was optimized and the interactions in the model and successful results were discussed. Finally the last chapter is the conclusion and present a short summery of the thesis. Here, the chapter reviews the achievements, and outcomes of the thesis. Possible future works also are discussed at the end of this chapter. The thesis also includes three appendixes in which appendix A and B covers two major side works that were carried out by the author of the thesis during the PhD study and appendix C presents the G-Code used to control the bioprinter’s robotic platform to print the brain model. 18 Chapter 2 Fabrication of 3D Brain Model 2.1. Defining the brain model The present bioprinted co-culture brain-like microstructure recapitulates the mechanical, micro-anisotropic, and synapse functional properties (i.e. synaptic transmission) of the brain (Figure 2.1). Figure 2.1. Schematic of the astrocytes and nerve fibers with neurons patterns in the brain. As a proof of concept, two major brain cell types (i.e., neural stem cell-derived neurons and mouse cortical astrocytes) were chosen in this research to create neuron-laden aligned nerve fibers in an astrocyte-laden matrix. The first is aimed to mimic the anisotropic arrangement of neuron cells in the aligned nerve fibers in the brain through embedded 3D bioprinting of a neuron-laden gelatin methacryloyl (GelMA)/gelatin/alginate/laminin bioink (Figure 2.2). Using the embedded bioprinting contributed to the gradual dissolution of gelatin at 37 ℃ during the culture and incubation 19 process [88]. Moreover, extension and migration of cells could occur in the alignment embedded bioprinted fibers concurrently.[89, 90] The objective of the astrocyte-laden GelMA matrix, with its relatively low Young’s modulus is to mimic the non-neuronal tissue parts in the brain by creating an astrocyte-laden GelMA hydrogel with relatively low Young’s modulus, which replicates the natural brain stiffness and enables growth and extension of astrocyte cells.[91]. In addition to providing brain-like stiffness, the astrocyte-laden GelMA hydrogel containing calcium ions acts as a supporting bath to facilitate the fabrication of the neuronal channels as well as improving the alignment of neurite growth in the bioprinted fibers (Figure 2.2). During the bioprinting process, the divalent calcium ions in the supporting bath interact with the α-(14)-linked chain configuration (i.e., egg-box structure) in the alginate chains, forming a physical crosslinking tool to fixate the printed fibers[92]. The printing process is then followed by a GelMA crosslinking step through UV light treatment. The variation in Young’s modulus between the printing fibers and the supporting bath ensures that the neuron cells are constrained to grow in the bioprinted fibers (Figure 2.2), thus, a brain-like micro-structure with aligned neuronal filaments surrounded by astrocytes can be achieved to study the neuron-astrocyte interactions as the neurites extend out from the neuron cells to form interconnected networks. The fabrication process of the embedded bioprinted brain-like microstructure is discussed in details in experimental procedure. 20 Figure 2.2. Schematic diagram denoting the fabrication process of 3D brain-like co-culture constructs using embedded printing. The neural cell-laden bioink is printed inside an astrocyte-laden support bath followed by UV crosslinking and in vitro culture. 21 2.2. Model and methodology I have developed a 3D brain model which could closely mimic the native brain tissue (Figure 2.3a). The 3D embedded printing configuration based on extrusion printing technology was used to fabricate the proposed 3D brain model construct. The model contains two types of cells (neuron cells and astrocytes) in which the filament structure of neurons are printed in a 3D structural shape. This models mimics the structure f neural fiber in native tissue (Figure 2.3a). Secondly, this model facilitated the study of the effect of cell-cell interaction as well as cell environment interaction in 3D (Figure 2.3b). The third feature of this model is that it facilitates the neuron alignment in the printed structure. The model consists of two main parts including the neural fibers and bulk support containing glial cells. The neural fiber consists of a composite hydrogel polymer of 4wt% GelMA/1wt% gelatin/2wt%Alginate/ 10 ug/ml laminin where neural cells are encapsulated inside the hydrogel. Alginate was used to enhance printability of the composite hydrogel solution and enable instant gelation of the composite polymer upon printing. The role of GelMA in the composite polymer is to facilitate the connectivity between the printed fiber and its surrounding and provide required mechanical properties. Laminin, on the other hand, provides the binding sites for neural stem cells to attach, so they can proliferate and differentiate into neural cells. Gelatin helps to make more porous material which is required for neuron cells to grow. The fibers are printed straight inside the supporting bath of GelMA that contains astrocytes. The supporting bath acts as the structural support for neural fibers and biological support for astrocytes (Figure 2.3c). After printing the fiber inside the bath of GelMA the entire construct is exposed to UV in which GelMA in the 22 bath and the fiber crosslink, forming mechanical-stable constructs that can be used for further study. Figure 2.3. a) The picture of a native brain structure. b) The schematic of the proposed brain model on chip consisting of two types of cells: astrocytes and neuron cells. c) The schematic of the fabricated 3D brain structure using an embedded printing technique. The model consists of the supporting bath including astrocytes and the printed fiber-like structure including neuron cells. d) The schematic of interaction of neuron-astrocytes in a 3D printed brain model. e) The schematic of alignment of the neuron cells in a fiber-like structure printed in a supporting bath. 23 Brain tissues consist of neuron fibers, where the fibers are surrounded by glial cells that supports neural functions. Here in the model, neuron stem cells are, therefore, printed in the form of fibers in the bath consist of astrocyte cells, mainly to mimic the native brain tissue structures. One advantage of this model is co-culture of two types of cells, required for studying the cell-cell interaction of astrocytes with neuron cells (Figure 2.3d). The neural stem cells are printed in the fibers and astrocytes are around the fibers creating the biological support for neural stem cells. In this model neural stem cells are differentiated to neurons using a neural growth factor (such as retinoic acid) and by tuning the mechanical properties of the fiber and surrounding constructs. Several factors will be optimized to facilitate differentiation of the neural stem cells into neurons including the composition of the construct, the mechanical property of structure, the density of neural stem cells and supporting astrocytes, the composition of the cell growth medium, and the growth factor. Another advantage of this model is the printed fiber-like structure which facilitates neuron alignments compared to the neurons randomly distributed in the bulk (Figure 2.3e). The alignment of neurons can direct axons and the neurite outgrowth from the neurons, facilitating better connectivity between neurons in the model. 2.3. Experimental procedure In the first step, printer head was calibrated and put in the home position, the x, y, and z axis were set according to the GCode of printing. Then, the syringe was mounted in the syringe pump and connected to the printer head (a 90 blunt syringe needle) with the plastic tube (0.38mm inner and 0.79mm outer diameters). In order to prevent leakage, the syringe was glued with the plastic tube. The flow rate of bioibnk was controlled by syringe 24 pumps (KD Scientific Inc., Holliston, MA, USA). For all experiments and the reliability of the results of each test, the syringe pump was continuously operated for 5 minutes with constant flow rate, and then the printing was started. Printing setup The 3D printing platform consisted of five parts including; 1) a syringe pump, 2) a syringe and tubing containing the printing ink, 3) a programmable robotic platform, 4) a printing head, and 5) printing bath. The printing setup is shown in Figure 2.4. Figure 2.4. The printing setup used for embedded printing of the brain model. Here, NE-1000 Programmable Single Syringe Pump was used to continuously pump the printing ink with volume rate of 4µL/min. The printing ink was loaded into a BD 1 mL Tuberculin Syringe and assembled on the syringe pump. The syringe was connected 25 to the printing head through tubing. A 30 gauge needle was used as the printing head. The connection between tubing, syringe and the printing head was arranged using a pair of needles in which one of them connect the syringe to the tubing and the other connect the tubing to the printing head. The cellink inkredible bioprinter was used as the programmable robotic platform. Note that it was possible to use the bioprinter to print the ink as well using its constant pressure mechanism. However, to achieve printed fibers with more consistent dimensions, I preferred to deploying the ink into the bath during the printing at constant flow rate (using a syringe pump) rather than the bioprinter constant pressure mechanism. Finally the last part of the printing setup is the printing bath which was made out of laser cut acrylic sheets. The bath contain 5%wt GelMA with cell media. The code was written in a way so that the printing head was periodically going into and out of the bath to print the fibers vertically. The G-code use for printing is attached in the Appendix C section. After setting up the bioprinter the bioprinter, the composite biomaterial for bath and bioink was prepared. For supporting bath solution, 5 wt% GelMA, 11 mM calcium chloride, and .2 wt% photo initiateor, 2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-propanone, (Irgacure 2959, Sigma Aldrich) were weighted and dissolved in 40% HBSS and 60% astrocyte complete media (DMEM+10% FBS+1% P/S. Then astrocytes with the concentration of of 8x10 6 cells/ml was suspended into the solution. Then the solution was incubated at 37 °C for 1 h. The bioink was prepared using 4 wt% GelMA, 1 wt% t gelatin and 1 wt%t Alginate (Alginic acid from brown algae, Sigma Aldrich) and 0.25% PI dissolved in 40% HBSS, 60% MEM and incubated at 80 °C for 50 minutes. Next, the ink was incubated at 37 °C for 30 minutes. Then, NE-4Cs were dissociated and dissolved in mixed with the bioink 26 supplemented with 10 ug/ml Laminin (Sigma Aldrich) at a concentration of 25x106 cells/ml. Polydimethylsiloxane (PDMS) molds were sterilized in 70% ethanol and UV overnight, followed by a thorough washing step with sterile DPBS. The supporting astrocyte-laden bath was pipetted into the molds and kept at 4 °C for 5-8 minutes prior to printing. Then, biomaterial of are placed in the PDMS molds which has dimensions of 1 cm 1 cm 0.5cm (Figure 2.5a). The mold containing the material was placed in the fridge (4C) for 10 minutes to physically crosslink the bath (Figure 2.5b). The bioink was loaded into 1ml syringes (BD) capped with a 30G needle (blunt, BD) and mounted on the extruder of a Cellink Inkredible bioprinter, and the tip of the needle was positioned at a designated point with respect to the position of the mold containing the supporting hydrogel, which represented the XYZ origin for all the printing configurations. The composite hydrogel fibers with neuron cells were printed inside the bath using an extrusion bioprinter and print the 3D structure of the hydrogel bioink by computer-assisted software. As it shown in the Figure 2.5c, the neural bioink consists of GelMA, alginate, gelatin, laminin and neuron cells which were pumped through the nozzle to form a fiber at the outlet in a CaCl2 bath. The calcium cations in the bath crosslinks in the ink, facilitate the formation of fibers. To print the neuron fiber inside the hydrogel bath, the fluids were pumped through the printing nozzle using syringe pumps to create the fibers. The speed and movement of the printing head is controlled by the printer, and the flow rate of the fluid is controlled by the syringe pumps. Finally, the printed structure are exposed to UV light (800 mW, 60 s) using an OmniCure S2000 machine for 30 second on each side in order to crosslink the GelMA material. 27 Next, the constructs were extracted from the molds, washed with DPBS and transferred to a well plate containing fresh MEM complete media (MEM+10% FBS+1%P/S). All printed constructs were incubated at 37 °C in 5% CO2 in well plates immediately after crosslinking. Culture media were changed the day after printing, and after that regularly every 2 days until staining at different time points. The fabrication mechanism consists of two crosslinking steps. First, ionic crosslinking of alginate in which alginate in the ink was crosslinked by the surrounded CaCl2 in the bath, providing partial structural stability of the fibers during the bioprinting process. Second, the entire scaffold was exposed to UV to crosslink GelMA material to create a solid structure where the fibers are connected through the GelMA network (Figure 2.5d). 28 Figure 2.5. The schematic of the bio-fabrication steps of the proposed 3D brain model. a) The cell laden hydrogel (Astrocyte cells) is dispensed in the PDMS mold. b) The construct is cooled down for 10 min. c) The bioink containing neuron cells are printed in supporting bath using extrusion printing. d) Photo-crosslinking of the construct is performed. 29 2.4. Material Preparation 2.4.1. Supporting bath hydrogel The supporting hydrogel was made by dissolving 5 wt% freeze dried GelMA, 11 mM calcium chloride (CaCl2), and 0.25 wt% PI 2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-propanone, (Irgacure 2959, Sigma Aldrich) in pbs followed by incubation at 37 °C for 1 h. 2.4.2. Cell suspended supporting bath The supporting hydrogel was developed suspending astrocytes at a concentration of 8x10 6 cells/ml in a DMEM-based solution containing 11 mM calcium chloride (CaCl2), 0.25 wt% PI 2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-propanone, (Irgacure 2959, Sigma Aldrich), and 5 wt% freeze dried GelMA. All components were dissolved in 40% HBSS and 60% astrocyte complete media (DMEM+10% FBS+1% P/S) followed by incubation at 37 °C for 1 h. 2.4.3. Bioinks The bioink was prepared using 4 wt% GelMA, 1 wt% t gelatin and 1 wt%t Alginate (Alginic acid from brown algae, Sigma Aldrich) and 0.25% PI dissolved in 40% HBSS, 60% MEM and incubated at 80 °C for 50 minutes. Next, the ink was incubated at 37 °C for 30 minutes. 30 2.4.4. Cell-Laden Bioinks The bioink was prepared using 4 wt% GelMA, 1 wt% t gelatin and 1 wt%t Alginate (Alginic acid from brown algae, Sigma Aldrich) and 0.25% PI dissolved in 40% HBSS, 60% MEM and incubated at 80 °C for 50 minutes. Next, the ink was incubated at 37 °C for 30 minutes. Then, NE-4Cs were dissociated and dissolved in mixed with the bioink supplemented with 10 ug/ml Laminin (Sigma Aldrich) at a concentration of 25x10 6 cells/ml. 2.4.5. Synthesis of GelMA (Gelatin methacrylate) Gelatin methacrylate (GelMA) was synthesized following a protocol mentioned in [88]. In short, 10wt.% gelatin from porcine skin (G1890, Sigma-Aldrich) was dissolved into Dulbecco's phosphate-buffered saline (DPBS) at 50C using a magnetic stirrer (240rpm). Methacrylic anhydride (MA) (Sigma Aldrich) was added to the solution drop wise, in order to reach a final concentration of 1.25wt.% MA. After 2 hours at 50C with constant stirring, the solution was diluted 1:1 with DPBS (previously warmed) and then dialyzed against distilled (DI) water with a dialysis membrane (Spectra/Por molecular porous membrane tubing, MWCO 12-14,000, Fisher Scientific) for 5 days at 40C, while stirring constantly (500rpm) and changing DI water every 12 h. Afterwards, the gel in the membrane is diluted two times and the final solution was filtered in 2 steps: firstly with a coffee filter, and second step with a sterile vacuum Express Plus (0.22 μm) Millipore filtration cup. The filtered solution was stored at -80C and, according to the experimental procedure, freeze-dried for 5 days. The methacrylation rate of GelMA was measured to be 40%. 31 2.5. Protocols and procedures 2.5.1. Cell cultures and neural stem cell differentiation Astrocytes were isolated from day 1 rats following a standard protocol.[93] Briefly, rat cortices were dissected, the hemispheres were cut into small pieces, incubated in Hank’s balanced salt solution (HBSS, Gibco) and 2.5% trypsin, and centrifuged to obtain a pellet of cortex tissue pieces. Through mechanical trituration a suspension of dissociated cells was obtained. The astrocytes were cultured in Dulbecco’s modified Eagle medium (DMEM, Gibco) supplemented with 10% fetal bovine serum (FBS, Gibco) and 1% Penicillin Streptomycin (P/S, Gibco). Neuroepithelial (NE-4C, CRL-2925, ATCC) cells were cultured and expanded in Minimum Essential Medium Eagle (MEM, Thermo Fisher Scientific) supplemented with 10% FBS and 1% P/S. To induce neuronal differentiation, 1 μM retinoic acid (RA) was added to the NE-4C culture for 48 hrs and the cells were cultured in MEM, 1% Insulin-Transferrin-Selenium (ITS, Thermo Fisher Scientific), 2 mM L-glutamine (Thermo Fisher Scientific), 1% P/S.[94-96] After this time, RA was removed and the culture was continued in Neurobasal medium supplemented with B27 (Gibco). For 2D differentiation, the wellplates were coated with Poly-D-lysine hydrobromide (PDL, Sigma Aldrich) prior to differentiation induction. All cell cultures were passaged according to the protocol provided by the respective vendors. 2.5.2. Embedded bioprinting procedures Polydimethylsiloxane (PDMS) molds were sterilized in 70% ethanol and UV overnight, followed by a thorough washing step with sterile DPBS. The supporting astrocyte-laden bath was pipetted into the molds and kept at 4 °C for 5-8 minutes prior to 32 printing. The bioink was loaded into 1ml syringes (BD) capped with a 30 gauge needle (blunt, BD) and mounted on the extruder of a Cellink Inkredible bioprinter, and the tip of the needle was positioned at a designated point with respect to the position of the mold containing the supporting hydrogel, which represented the XYZ origin for all the printing configurations. A syringe pump (New Era Pump Systems) was used to extrude the bioink at 3 μl/min rate. After printing, the full structures were crosslinked under UV light (800 mW, 60 s) using an OmniCure S2000 machine. Next, the constructs were extracted from the molds, washed with DPBS and transferred to a wellplate containing fresh MEM complete media (MEM+10% FBS+1%P/S). All printed constructs were incubated at 37 °C in 5% CO2 in well plates immediately after crosslinking. Culture media were changed the day after printing, and after that regularly every 2 days until staining at different timepoints. 2.5.3. Viability and metabolic activity assays The cell viability inside the hydrogel was assessed up until the day in vitro (DIV) 14 through a calcein-AM/ethidium homodimer Live/Dead assay (Invitrogen) and imaging with an inverted fluorescence microscope (Zeiss Axio Observer D1 Fluorescence Microscope, Carl Zeiss, Germany). The cellular metabolic activity throughout the culture lifespan was assessed using Presto Blue Reagent (Life Technologies). After 2 hours of incubation at 37C, the reagent was removed and the absorbance was measured using a microplate reader (BioTek Synergy 2, Vermont, USA), normalizing the extracted values to the 600 nm reference wavelength. The metabolic activity was monitored on DIV 1, 3, 7, and 14. 33 2.5.4. F-actin/dapi staining The samples are washed and fixed by treating with 4% formaldehyde for 20 min at room temperature and the exceed formaldehyde is washed out again 3 times with DPBS. In the next step, the cell membranes are permeabilised with 0.1% Triton X-100, C14H22O(C2H4O)n, for 10 minutes and washed again 3 times with DPBS. For F-Actin and DAPI staining, a 1:40 phalloidin, conjugated with Alexa Fluor 488 dye, in DPBS solution is prepared and dropped on top of the biological sample and incubated for 45 minutes. The excess dye is removed by washing the samples 3 times with DPBS, intervals of 10 minutes between every washing. Last but not least, the DAPI, diluted in DPBS (1:1000), is added and the samples are incubated for 10 more minutes. DAPI (4’,6-Diamidino-2-phenylindole) is used to stain DNA. The blue fluorescent nucleic acid binds to double-stranded DNA and the fluorescence is increased 20-fold. The excitation and emission wavelength for the DNA-DAPI complex is 364 nm and 454 nm, respectively. To label filamentous actin (F-Actin) the Alexa Fluor 488 dye is conjugated to phalloidin. Phalloidin is a seven amino acid peptide toxin, which binds with high affinity and sensitivity to polymerized actin. The excitation wavelength is 495 nm, whereas the emission spectra is around 518 nm. 2.5.5. Immunochemistry The immunocytochemical characterization was performed by fixing the cultured cells in the structures of 4% paraformaldehyde (PFA, Electron Microscopy Sciences) in DPBS for 20 min, treating them with 0.1% Triton-X 100 (Sigma-Aldrich) for 10 min, and washing with DPBS 3 times. After fixing, cells were incubated with primary antibodies 34 diluted in PBS and 10% bovine serum albumin (BSA) overnight at 4°C. The primary antibodies used were the mouse anti-beta-III tubulin monoclonal antibody (anti-β III tubulin Abcam) and the rabbit anti-glial fibrillary acidic protein (anti-GFAP, Abcam) monoclonal antibody at a dilution of 1:200. Conjugated secondary antibodies (goat anti-mouse IgG, Invitrogen Alexa Fluor 488, and goat anti-rabbit IgG, Invitrogen Alexa Fluor 594, Thermo Fisher) and DAPI counterstaining were used for visualization under the fluorescence (Zeiss Axio Observer D1) and confocal (SP7, Leica) microscopes. Z-stack confocal imaging was performed on a Zeiss LSM 880 airyscan microscope. All images were analyzed using Fiji and ImageJ softwares. 2.5.6. Statistical analysis The statistical analysis was performed using the Graphpad Prism software. The data was evaluated using one-way ANOVA followed by post hoc test (Duncan’s test) with p < 0.05 considered significant. 2.5.7. UV Crosslinking The gelatin macromers containing MA pendant groups can be crosslinked with the photoinitiator Irgacure 2959 under a UV light. The MA groups react to create a hydrogel. The 200 Watt UV lamp can be calibrated using a radiometer. For this project the intensity is set to 850 mW with a distance to the probe of 8 cm. This setting results in an irradiance of 6.9mW/cm2.The time for UV curing can be adjusted at the control panel of the OmniCure S2000. 35 2.5.8. Mechanical testing The compressive modulus of 5% chemical GelMA hydrogel was tested at a rate of 20% strain per min on an Instron 5542 mechanical tester (Instron, 5942 Single Column Tabletop Model, an ITW company (Illinois Tool Works Incorporation, Glenview, Illinois, USA). The software package used is Bluehill (Version 3) provided by Instron. The compressive modulus was evaluated as the slope of the linear region corresponding with 0–5% strain. 2.5.9. Rheological Characterization To evaluate the rheological properties, material responses to shear were determined in continuous flow experiments with a linearly ramped shear rate from 1 to 100 s− 1. The baths and bioinks responses to the application and removal of shear were examined in shear-recovery and printability experiments, where oscillatory (10 Hz) time sweeps were conducted with alternating high/low strains of 0.5/250% every 2 min. 36 Chapter 3 Development of Printing Biomaterial 3.1. Development of 3D bioprinting material The two key GelMA properties, physical gelation and photo-crosslinkability, facilitate the embedded printing process for creating co-culture brain like microstructures. The reversible physical crosslinking of the GelMA bath at 4℃ creates sufficient stiffness in the bath to support fiber extrusion during printing. Further, the alginate element in the ink is crosslinked by the calcium ions dispersed in the bath thus creating well-defined fibers with crosslinked boundaries. Calcium chloride is added to the bath to provide a physical crosslinking for the gelatin/alginate/GelMA/laminin bioink. [97] The printed construct undergoes a chemical GelMA crosslinking through exposure to UV radiation, and therefore any unexpected dissolution of GelMA polymers at the 37 C is prevented during the culture and incubation process. In order to create a support for embedded bioprinting process and then print neuron fiber into the supporting bath; several parameters of supporting bath and neuron fiber need to be optimized separately. 3.2. Optimizing Supporting bath To make a desired supporting bath for the embedded bioprinting process, I have optimized the parameters of GelMA supporting bath including the concentration and ratio of material components, rheological properties, mechanical modulus, UV curing time, and degradation behaviors of GelMA hydrogel. In the first step, the concentration of GelMA used for the supporting bath needs to be optimized. In order to find the optimal concentration of GelMA, two tests of crosslinking and young modulus were carried out. By systematic test of UV crosslinking 37 time and different concentration of GelMA the optimal UV exposure time and GelMA concentration to make a gel were identifies. For this test, the concentration of photo initiator was kept constant (i.e., as 0.25 wt%.). To find out the optimal UV exposure condition, a fixed concentration of the photoinitiator (0.25% w/v) was used, and hence various concentrations of GelMA solution (from 2.0% w/v to 5.0% w/v) and exposure time of UV light (from 60 seconds to 240 seconds) were tested systematically to study chemical formation of GelMA hydrogel. As shown in Figure 3.1a, the GelMA hydrogel with the concentration of less than 3 wt.% could not be crosslinked. This suggests that the polymer chains in the low concentration of GelMA solutions provide a low crosslink-density resulting in the occurrence of the poorly-crosslinked structure in chemical GelMA hydrogels [98]. Since the UV time exposure needs to be minimized with the existence of cells (with high cell viability) the UV crosslinking time was picked for 60 s. With this condition, the concentration of GelMA higher than 3 wt.% (3-5 wt.%) were tested to find the concentration of fully crosslinked GelMA at 60 s exposure time. As shown in Figure 3.1b, 4.5 wt.% and 5 wt.% were the concentrations of GelMA which were fully crosslinked at the 60s exposure time. In addition to the crosslinkability test, the modulus of the chemical GelMA hydrogels after the exposure to UV was determined. It is reported that the compressive modulus of a native brain tissue is in the range of 0.6 kPa to 1.2 kPa [99]. In order to find out the optimal GelMA concentration for the supporting bath which is in the range of the reported mechanical property of the native brain issue, I have measured the compressive (young) modulus of 4.5 wt.% and 5 wt.% which were crosslinked at 60 s UV exposure time. The compressive modulus of 5% chemical GelMA hydrogel tested using Instron 38 Machine was found to be 0.71 ± 0.10 kPa (Figure 3.1c), like the stiffness of brain tissues. On the other hand, I also evaluated the effect of the UV exposure time on the modulus of 5 wt.% chemical GelMA hydrogels. I found that by increasing the UV exposure time the comprehensive modulus is increasing. The young modulus of 5wt.% GelMA exposed to UV for 70s is 1 kPa which is still located in the range of the brain modulus while the modulus of 5 wt.% GelMA with the exposer time larger than 80s is over than 1.5 kPa which much higher than the modulus of the brain tissue (Figure 3.1d). The minimum requirements for a material to be used as a supporting bath in embedded bioprinting is to follow shear-thinning behavior as well as self-healing properties. These properties are necessary to provide a fast-recovery behavior for fixing the embedded printed structures. Various types of physical interaction, such as electrostatic hydrophobic forces and hydrogen bonding exist in the GelMA hydrogel, which contribute to the viscoelasticity of the final hydrogel composition (Figure 3.2 (i)). Upon insertion of the needle into the GelMA bath, some of the electrostatic and hydrogen bonds disentangle due to the imposed shear stress by the needle movement (Figure 3.2 (ii)). This behavior creates aligned polymer chains and thus lower viscosity while printing. After the needle (i.e, shear stress) removal, the physical bonds are formed again and the GelMA prepolymer chains returns to the original randomly oriented configuration, a property called self-healing behavior. 39 Figure 3.1. a) The chemical GelMA hydrogel formation from the GelMA solutions (with different concentrations) by treatment with various UV curing times. It was found that under a low UV intensity (800 mW), the 60s exposure period is the minimal time for obtaining a fully-crosslinked chemical GelMA gels from 3-5 wt.% GelMA solution. b) The crosslinked morphologies of GelMA hydrogels with various concentrations after exposure to UV (for 60 s) from top and bottom sides. c) The compressive modulus testing results for the chemical GelMA hydrogels. d) The compressive modulus of 5 wt.% chemical GelMA hydrogel exposed to UV for 70s, 80s, and 90s. 40 Figure 3.2. Schematic process of the dynamic structure change of the GelMA solution under a shear tress. Increasing the GelMA concentration- thus the internal physical interactions- creates a significant elevation in the hydrogel viscosity (Figure 3.3). Moreover, a previous study reported that presence of calcium ions (>10 mM) generated an electrostatic interaction which affected the viscoelastic properties of gelatin hydrogel.[100] The effect of the presence of Ca2+ on viscosity was found to be prominent, compared to the case where calcium ions are absent, only when the concentration of GelMA is below 3 wt% (Figure 3.3 and 3.4). Notably, no significant difference in viscosity could be observed by adding Ca2+ to 5 wt% GelMA hydrogels, indicating that addition of Ca2+ would not affect the shear-thinning behavior in 5 wt% GelMA concentrations. Based on these results, it is suggested that calcium ions selectively form a complex with carboxyl groups on the low 41 GelMA concentration backbone (i.e., ─COO-─Ca2+─ COO-─ bond), which links the adjacent GelMA chains, forming an internal network and increasing the viscosity of GelMA hydrogels. In high GelMA concentrations, however, the effect of the secondary forces such as the hydrogen bonding and hydrophobic interactions[101] among the GelMA chains are stronger than those imposed by Ca2+. As such, the viscosity in 5 wt% GelMA hydrogel is not impacted substantially by the presence of calcium ions.[100] Figure 3.3. The effect of calcium ions on the viscosity of the GelMA solution with the different concentrations. 42 Figure 3.4. The effect of calcium ions on the shear-thinning behavior of 5% GelMA baths. Moreover, another characterization is the degradation test since the GelMA material is placed in an incubator (after printing) for at least 14 days for cell culturing. Therefore, the crosslinking hydrogels must be stabilized at physiological environment upon incubation at 37 0C. Insufficient crosslinking can lead to expedited and undesired hydrogel degradation during the incubation process before cell-laden construct maturation. Therefore, I tested the degradation behavior of chemical GelMA hydrogels with various UV exposure times. A slight weight loss has been observed after 7 days of incubation of chemical GelMA hydrogels exposed to the UV light for 60, 70, and 80 (Figure 3.5). However, after another 7-day of incubation, the chemical GelMA hydrogels in these groups had no significant weight loss (Figure 3.5), suggesting that an undesired small degree of dissolution or degradation in uncrosslinked GelMA polymers which might cause the weight loss of the hydrogel in the medium at the physiological temperature. 43 Figure 3.5. The degradation behaviors of 5 wt.% chemical GelMA supporting hydrogel after 7 and 14 days of incubation. In terms of the embedded 3D bioprinting technique, the supporting bath requires certain rheological properties (e.g., self-healing property). In order to characterize the self-healing property, the strain-induced damage and healing of 5 wt.% GelMA hydrogel with calcium ions was evaluated through a continuous oscillatory strain sweep [102] (between 250% and 1% strain at 10Hz). An application of 250% strain was performed for 2 min to destroy the structure of hydrogel; subsequently, the hydrogel was allowed recovery by decreasing the strain to 1% for another 2 mins. Under the high dynamic strain of 300% (Figure 3.6a), the storage modulus value (G’) of GelMA hydrogel was decreased to lower than the loss modulus (G”), and hydrogel showed a gel-like-to-sol transition behavior. In contrast, when the strain moved down to 1%, the G’ of GelMA returned to the initial value, which showed the occurrence of a rapid sol-to-gel-like transition. In other words, the rapid hydrogel transition from the fluid-like or gel-like behavior indicates that the GelMA hydrogel has a self-healing ability to restore the destroyed structure after the movement of Time (s)60s 70s 80s 90sWeight Percentage (%)0501001502000 day 7 days 14 days 44 the printed nozzle. In Figure 3.6b, it is also shown that 5wt.% GelMA as a supporting bath is cracked by the movement of the printed nozzle right after printing (obvious after crosslinking). However, after incubation of 5 wt.% GelMA at 37C, the cracked part is recovering, and after 24 hours GelMA is fully healed. Therefore, these results have demonstrated that 5% GelMA hydrogel possess the self-healing behaviors which is the basic required rheological behavior of a supporting bath for embedded printing. Thus, it is used here for the brain-like structures. A point to consider in embedded bioprinting is the creation of cracks inside the bath upon the traveling of the needle during the printing process. Such cracks could lead to unconfined cell growth, undesired changes in structure stiffness and compromised shape quality. [97] The reversible transition from fluid-like to gel-like state provides GelMA with a self-healing capability to restore its uniform structure throughout the printing process and upon crack formation (Figure 3.7a). To confirm the self-healing capability of the crosslinked GelMA hydrogels, Scanning Electron Microscopy (SEM) imaging was performed on the samples in three stages (Figure 3.7b). First, the porous structure of the GelMA hydrogel was imaged by SEM prior to the printing process (Figure 3.7b (i)). Next, the bioink was printed inside the bath at room temperature, and the crosslinked construct was imaged in SEM to confirm the creation of crack in the bath (Figure 3.7b (ii)). Upon incubating the printed and crosslinked construct at 37 C, the SEM images confirmed that the crack was healed to a great extent (Figure 3.7b (iii)). Consequently, with their shear-thinning and self-healing properties, 5 wt% GelMA hydrogels can satisfy the rheological requisites of a support bath for embedded printing. 45 Figure 3.6. a) Oscillatory strain sweep between low (1%) and high (250%) strain shows 5% physical GelMA hydrogel with a rapid self-healing ability during the printing process. b) The scratch recovery of 5% GelMA after the movement of the nozzle, showing that the scratch was recovered within 30 min. 46 Figure 3.7. (a) Schematic process showing the self-healing behavior of GelMA baths under shear stress triggered upon the needle movement. (b) SEM images of 5% GelMA bath before (left) and after (middle) the movement of nozzle in the bath showing the self-healing process (right) of the cracks in the bath 47 3.3. Optimizing bioink and 3D printing parameter Printing fiber-like structures and hierarchical networks in the supporting bath requires an injectable bioink capable of easy extrusion with no post-print deformation. Such injectable behavior of the bioink is defined by the rheological properties of the biomaterial.[103] Here, a combination of GelMA, gelatin and alginate (GM/Gel/Alg, GGA bioink) (Figure 3.8) was chosen to create fibers with tunable viscosity, neuron-compatible stiffness and porosity. Figure 3.8. Schematic diagram of the embedded bioprinted GelMA-based bioink As for optimizing the bioink, the rheological properties of the bioink reflect the injectable behavior of biomaterials [104]. A bioink used as fibers in the supporting bath requires to be extruded from the printed nozzle without any deformations after printing. In this research, the rheological properties of the gelatin/alginate/GelMA mixed bioink were 48 evaluated. First, the rheological properties were evaluated and optimized using composite bioinks with different GelMA concentrations in the GGA bioink (Figure 3.9). Figure 3.9a also shows a downward trend in the viscosity of the bioink occurred when the shear rate was increased using a continuous flow experiments similar to GelMA bath. These results demonstrate the shear-thinning property of the mixture of gelatin and alginate. Moreover, I also evaluated the solid-like to fluid-like transitional behavior by performing a continuous oscillatory strain sweep under alternating high/low strains (250% and 1%). Within the period of high strain treatment, the value of G”(loss modulus) surpassed that of G’(storage modulus), indicating that a rapid phase transition from solid to fluid at the high strain (Figure 3.9b). In contrast, under the low strain, the values of G’ and G” showed a mechanical recovery behavior, meaning the phase change from fluid to solid occurred in the bioink. This shear-thinning behavior and the rapid repeatable phase transition prove that the gelatin/alginate mixed bioink is a great candidate for extruding the ink from the nozzle. Notably, the continuous flow experiment and the continuous oscillatory strain sweep were used for evaluating the shear-thinning and phase transition properties for both the supporting bath and bioink. From the rheological point of view, the interpretation of the results is different between the supporting bath and the bioink: in terms of the embedded printing technique, the shear-thinning and phase transition properties for the bioinks represent their printability; and these properties for the bath enable (i) the free movement of the nozzle in the supporting bath, and (ii) rapid restoration of the destroyed the printed structures after the movement of the nozzle [105]. 49 Figure 3.9. a) Continuous flow of 2% alginate mixed with 7%, 5%, 3.5%, and 3% gelatin tested as bioinks at an equivalent shear rate. The results show that the bioinks have the shear-thinning behavior. b) Oscillatory strain sweep between low (1%) and high (250%) strain shows the rapid phase transition from a gel-like to fluid-like behavior for the 5% gelatin/2% alginate bioink, i.e., necessary for localization of extruded bioinks. 50 As it shown in the Figure 3.10, a declining trend occurred in the viscosity upon shear rate increment at a continuous flow, demonstrating the shear-thinning characteristics of the bioinks. Moreover, in Figure 3.10, oscillatory stress model shows the storage modulus (G’) of bath-5% is higher than other inks conditions, it suggests that the 5% GelMA as a bath is with ability to maintain the fiber structure after printing. 3, 4, and 5% inks represent 3, 4 and 5%GelMA in the inks. All inks include 1% gelatin, and 2% alginate. Figure 3.10. Oscillatory stress model shows the storage modulus (G’) of bath-5% is higher than other inks conditions, it suggests that the 5% GelMA as a bath is with ability to maintain the fiber structure after printing. 3, 4, and 5% inks represent 3, 4 and 5%GelMA in the inks. All inks include 1% gelatin, and 2% alginate. 0.1 1 10 100 10001101001000 G' (Pa)Stress (Pa) Bath 5% Ink 3% Ink 4% Ink 5%51 The ink extrusion parameters, such as printed fiber diameter, movement speed of the nozzle and the pump flow rate regulate the thickness of printed fibers and the quality of printed structures.[106] the mentioned parameters were optimized by printing the bioinks in the supporting bath. The parameters of extrusion (such as the diameter, movement speed of nozzle and flow rate) need to be characterized. Previous studies have reported that the nerve fiber tracts in a brain (i.e., association, commissural, and projection fibers) consist of axon populations of neurons [107], which have diameters more than 200 μm [108]. Therefore, a 27G and 30G needles (with the inner diameters of 210 μm and 160 μm, respectively) were used in this study for printing fibers under various movement of nozzle speeds (mm/s) and extrusion rates (μL/min). Generally, the thickness of printed fibers is also associated with the nozzle speed and the extrusion rate. As shown in Figure 3.11, by decreasing the flow rate of the bioink flow, the printed fiber diameters decrease at a fixed printer head speed. It is also shown that at a fixed flow rate an increase in the printer head speed reduces the fiber diameter. According to Figure 3.11, the average diameters of printed fibers of 30G nozzle at the various nozzle speeds of 2 mm/s, 6 mm/s, and 10 mm/s are 293.96 ± 6.31 μm, 223.00 ± 6.13 μm, and 142.43 ± 5.20 μm, respectively. It is also a well-known fact that the printed fibers are not continuous when the nozzle speed is too high or the extrusion rate is too slow. For example (see Figure 3.12), at the 3μL/min extrusion rate, the printed fibers are not continuous for the nozzle speeds of 6 or 10 mm/s. In contrast, a continuous printed fiber can be observed at the 3 μL/min extrusion rate and 2 mm/s nozzle speed. Thus, for printing an ideal fiber using 3D embedded bioprinting one needs to make sure that (i) under the fixed nozzle gauge, the fiber diameter is close to the nozzle gauge, and hence the shape 52 of the fibers was circular; and (ii) the nozzle moving speed is regulated to obtain the desired printed fibers with diameters expected. Figure 3.11. a) The diameters of fibers were printed from 27G nozzle at various nozzle moving speeds and extrusion rates. b) The diameters of printed fibers were extruded from 30G nozzle in the supporting gel under various nozzle moving speeds and extrusion rates. 53 Figure 3.12. The morphologies of fibers were printed from the a) 27G and b) 30G nozzles at the various extrusion rates and fixed nozzle moving speeds. 54 Figure 3.13. The circularity of printed fibers extruded from 30G nozzle in the supporting gel under various nozzle moving speeds and extrusion rates. Moreover, the cross-sectional geometry of printed fibers tends to form a circle in 3 μL/min extrusion rate and 2 mm/s nozzle speed (Figure 3.13). Previous studies indicated that the cells can maintain high viability under low extrusion rates (<5μL/min) during the printing process and a one-stroke painting spiral structure can be easily printed in the supporting gel.[90] Therefore, this configuration selected as the optimized condition to print well aligned straight fibers with circular cross-section inside the supporting GelMA bath (Figure 3.14). 3 5 10|0.00.20.40.60.81.01.21.4 CircularityExtrusion rate (l/min) 2 mm/s 6 mm/s 10 mm/s55 Figure 3.14. Phase contrast imaging of the extruded parallel fibers with circular cross-section and approximately 300 μm thickness The young modulus of different GelMA ink concentrations (3, 4, and 5 wt%) were measured (Figure 3.15). All inks include 1 wt% gelatin and 1 wt% alginate. By increasing the GelMA concentration from 3 wt% to 5 wt%, the young modulus is increasing from 0.4 kPa to 1.1 kPa. The brain stiffness previously reported is in the range of 0.6 kPa to 1.1 kPa. The ink with 4 wt% GelMA concentration, 1 wt% gelatin, and 1 wt% alginate with the 0.7 kPa stiffness was chosen for our work which fits in the brain stiffness range (Figure 3.15). 56 Figure 3.15. Young’s modulus of different GelMA ink concentrations. 3, 4, and 5% inks represent 3, 4 and 5%GelMA in the inks. All inks include 1% gelatin, and 2% alginate. The stiffness of our optimized condition (ink 4%) is in the range of brain stiffness. 57 Chapter 4 Cell Study of Developed Biomaterials 4.1. Cell study of bath and printed cell-laden construct Several physical and mechanical characterization of the supporting bath hydrogel and composite bioink were optimized for gelation, and printability to enable the proposed embedded bioprinting approach. In this section, the cellular responses to biomaterials developed in previous sections will be studied to optimize the biological responses of neural cells within these hydrogels. 4.2. Optimizing supporting bath for cellular activity of astrocyte The 3D printed brain model contains the fundament units of brain which is integrated with astrocytes, neurons, and neurites. The construct in this model were created by a supporting hydrogel bath with astrocyte cells and an embedded 3D neuronal network laden with neuron cells. In order to optimize the supporting bath for astrocytes to be alive, proliferate, and function, a series of experiments were carried out. First, a thin layer of the supporting gel (1 mm thickness) was used to test the astrocytes’ spreading ability in different conditions of the supporting gel (e.g., wt.% GelMA, UV exposure time, UV power density, and distance). Different concentrations of GelMA (3.5 wt.%, 5 wt.%, and 7 wt.%) were used, and astrocytes mixed with each concentration of GelMA were exposed to UV at different times (50s , 60s, and 80s). As shown is Figure 4.1, the optimized condition for astrocytes spreading was found to be 5 wt.% GelMA, 60s exposure time, 800 mW, and at a 5-cm distance. 58 Figure 4.1. The morphologies of astrocytes in the 3.5%, 5%, and 7% chemical GelMA gel with different UV exposed times after 14 days of incubation. It is shown that the astrocytes could extend and have spindle morphologies in 5% chemical GelMA hydrogel under the 60s UV curing time. 59 To further assess the viability of astrocytes within in various UV crosslinking times, a live and dead assay was performed on the constructs after UV crosslinking (Figure 4.2). High cell viability was observed for exposure times less than 90sec. The cell viability significantly diminishes for exposure time higher than 100 sec. Figure 4.2. Live/dead viability assay of astrocytes in 5% GelMA bath upon different UV exposure treatments. Astrocytes encapsulated in a thin layer of 5wt% GelMA bath were also labeled with an astrocyte specific marker (GFAP) to study cell morphology and network formation in the bath (Figure 4.3). Confocal imaging and 3D reconstruction of the immunostained constructs on day 7 (Figure 4.3a) revealed that most astrocytes started to extend into a spindle morphology and interconnected. Astrocytes spreaded more into the bath and a larger network of astrocytes were observed on day 14 (Figure 4.3b). 60 Figure 4.3. GFAP/ DAPI staining of astrocytes in the 5wt% chemical GelMA supporting gel after a) 7 days of incubation, and b)14 days of incubation. As the printed model has a higher depth with the thickness of 5 mm, the above-mentioned optimum condition was used to develop the final supporting gel structure with a thickness of 5 mm. However, the increase of the thickness brought forward the need of tuning the previous conditions due to two main issues: (i) the need of providing nutrient 61 equally to all the sections of the structure, and (ii) finding a proper balance between the gel resistance to degradation and its ability to sustain cell proliferation. For the first issue derived from the increased thickness (resulted in the deficiency of nutrients in the most inner parts of the supporting gel), the live/dead assay tests have shown that the center of the structure has a much higher number of dead cells compared to the outer parts. In order to overcome this issue, couple of pillars were placed in the supporting gel during crosslinking in order to generate hollow channels that could allow for a better perfusion of nutrients (Figure 4.4). Successive live/dead assay tests proved this approach to be successful: see Figure 4.5 that shows that the viability of the cells in the inner parts increased considerably from the structure without pillars to the one with posts. Figure 4.4. Schematic of the crosslinked supporting gel containing astrocytes: a) the structure without pillars, b) the structure with posts. 62 Figure 4.5. Live/dead assay for MCF-7 at day 7: a) live/dead assay for the structure with pillars, b) live/dead assay for the structure without pillars. The images were taken by a florescent microscope. The blur part shows that the cells are in different planes of focus. The second issue required tuning of the exposure time in order to meet the tradeoff between supporting gel degradation and cell viability and spreading. The previous optimum exposure time did not allow for the full crosslinking of the structure. On the other hand, an excessive increase would hinder the cell spreading ability in the supporting gel. In order to overcome these issues, a fine-tuning on the new conditions had to be carried out. A set of new exposure times and cell concentrations was tested, using cell types that would allow for a proper modeling of astrocytes: 3T3 cells (fibroblast cells) were used to test the cell spreading ability, while MCF-7 cells (breast cancer cells) were used to test the supporting gel’s stiffness as suitable to host astrocytes. Although 3T3 cells show optimal spreading on stiff substrates with micro- or nano-scale roughness (e.g., Flasks), MCF-7 cells were chosen as they are much softer, like in the case of astrocytes. In this framework, 3T3 spreading would mimic the spreading behavior of astrocytes, whilst MCF-7 63 aggregating in colonies would signify that the stiffness of the supporting gel is also appropriate for astrocytes. These tests led to the optimal condition of 70s exposure time, 800 mW, 5 cm distance with a concentration of 8 million/ml cells. These optimal conditions were finally tested for spreading of 3T3 (as shown in Figure 4.6). The results show that 3T3 cells were successfully spreading in these conditions. Figure 4.6. Actin/ DAPI staining of 3T3 cells at different days: a) day1, b) day2, c) day3. The images were taken by a florescent microscope. The blur part shows that the cells are in different plane focus. After optimizing the GelMA concentration, UV exposure time, and UV intensity for the supporting bath as well as incorporating pillars to make hollow channels for transferring nutrients to the cells close to the center of the supporting bath, astrocytes were encapsulated into optimized supporting bath. I further tested the viability after day 1 for astrocytes. Figure 4.7a shows a good viability results for astrocytes encapsulated in 5 wt.% GelMA. The viability results also carried out for 5 wt.% GelMA with different UV exposure times. As shown in Figure 4.7b, an increase in the UV exposure time decreases the viability of cells. The metabolic activity of astrocytes is also measured for different days (day 1, day 7, and day 14). The astrocytes did not proliferate in the structure (Figure 64 4.7c). The reason is the astrocytes were used here had the passage number of 5, which is the maximum passage number that astrocytes isolated from a rat can proliferate. After performing the cell viability and metabolic activity on astrocytes, astrocyte cells were labeled by an astrocyte specific marker (GFAP) for identifying their morphologies and the network formation in the supporting bath. After 7 days of incubation, the results of confocal 3D reconstruction images show that most astrocyte cells kept their round shape; only a few astrocyte cells extended with a spindle morphology and connected with other cells (Figure 4.8a). Notably, the proliferation behavior and the entire network formation of astrocyte cells can be observed after 14 days of incubation due to the degradation of the supporting bath (Figure 4.8b), enabling astrocytes with the free spatial space to grow and spread throughout the supporting bath. 65 Figure 4.7. a) The live/dead assay and cell viability of astrocytes in the 5% chemical GelMA supporting gel after UV treatment. b) The cell viability percentage of astrocytes at different UV exposure times. c) The metabolic activity of astrocytes for different UV exposure times at three different time points (1 day, 7 days, and 14 day). 66 Figure 4.8. GFAP/ DAPI staining of astrocytes in the 5wt% chemical GelMA supporting gel after a) 7 days of incubation, and b) 14 days of incubation. 4.3. Bioink Optimization for cellular activity of neuron cells After optimizing the supporting bath, the bioink material laden neuron cells has to be characterized. In order to optimize the ink, a series of experiments with different ink materials was carried out. For optimizing the ink condition, 3T3 cells used to characterize the ink conditions. After optimizing the ink with 3T3, neuronal cells will be used with the optimized ink for further experiments. Different conditions of the gel (e.g., wt% Gelatin, GelMA, UV exposure time, UV power density, and distance) were tested to optimize the 67 ink for 3T3 spreading and viability. The results showed that the optimized conditions for 3T3 spreading are 5 wt.% GelMA, 2 wt.% alginate, 70s exposure time, 800 mW power, and 5 cm distance. Figure 4.9 shows the confocal image of 3T3 cells spreading after 14 days. The images validate that the proposed bioink used for the fiber structures is suitable for the cell growth. Moreover, the fiber structure caused the 3T3 cell to mainly grow in the direction of the fibers. It is also interesting that the printed fibers improved the alignment of the actin filaments in 3T3 cells and caused the cells to elongate in the fiber direction. This study suggests that the printed fiber structures could direct the growth of the cells and improve their alignment inside the construct. Therefore, it was hypothesized such structure can facilitate the growth of the neural cells in the direction of fibers and improve their alignments. This could result in better connectivity between neurons. 68 Figure 4.9. Actin/ DAPI staining of 3T3 cells encapsulated in the GelMA/alginate ink and treated at a) 70 s, and b) 80 s UV after 14 days of incubation. Next, to build neuron-laden fibers in the brain-like construct, the neural stem cells (NE-4Cs) in the GM/Gel/Alg bioink were encapsulated and printed parallel structures inside the bath using embedded bioprinting (Figure 2.2). Upon extrusion of the bioink in 69 the bath, the Ca2+ ions present in the surrounding bath crosslinked the alginate chains around the fibers to create a well-defined patterned structure and prevent the fibers from diffusion before UV crosslinking. At the first step to evaluate the process of printing of neural stem cells encapsulated in the bioink in the supporting bath, live/dead assay was used to stain live and dead cells of neural stem cells right after printing (Figure 4.10). As shown in the figure 4.10, most of cells were alive and the printing process were optimized with this type of cells. Figure 4.10. (a) Phase contrast images and (b)confocal images of live/dead staining of neural stem cells in the printed fibers after embedded printing in the 5% GelMA bath. To further study the morphology of printed neural stem cells in the structure, actin/dapi staining of printed neural cells at day 7 were performed (Figure 4.11). As shown in the figure, the neural stem cells were alighned in the fiber structure and they did not diffuse to the bath and fabricated alighned fiber structure in the bath after 7 days. 70 Figure 4.11. Phase contrast images of (a) morphology and (b) F-actin/DAPI staining of neuron cells in the printed fibers after embedded printing in the 5% GelMA bath. Confluent parallel fibers with a thickness of around 300 μm are created upon 7 days of in vitro culture. 4.4. Summary In this chapter, the printining bath condition was optimized to achived high cell viability as well as robust astrocyte growth. Several parameters including mechanical properties, gel composition and cell media was optimizied to achieve robust growth of acrocytes in the bath. 71 Chapter 5 Neural Stem Cell Differentiation 5.1. Neural stem cell differentiation into neuron Neuron stem cells from the rat (neuroepithelial, NE-4C) were cultured and differentiated into neurons for the proposed study. Using the stem cells and differentiating them to a specific cells is cost effective and provides the potential ability to use human stem cells for future work. 5.2. Stem cell differentiation in 2D Differentiation of NE-4Cs was optimized in 2D following the previous 2D-based protocol [109] to differentiate neural stem cells into neurons in 2D (Figure 5.1a). First, confluent NE-4C cells (with the passage not more than 5) were exposed to retinoic acid for 48 hours, and the media of the cells was changed. Retinoic Acid (RA) was shown to induce neuronal differentiation in NE-4C subcultures [109]. Upon addition of RA, neural stem cells undergo morphological changes, aggregation and colony formation, and start to show neuronal characteristics within 12 days and in five stages. In the first stage, the cells start to aggregate and form colonies after 2 to 3 days of induction. In the second phase, the cells start to produce axon elongation on days 4 to 5. After that the cells start to elongate all through the surface. This stage of differentiation observed at a 2D surface. The differentiated cells in 2D were then mixed with the bioink material to utilize them in 3D. However, elongation of cells was not observed in 3D. To induce the differentiation, I exposes them to retinoic acid (RA). However, elongation of the neuron cells did not occur. Then, I decided to differentiate NE-4Cs in the bioink instead of differentiating in 2D and using the differentiated cells in 3D. The NE-4Cs were mixed with the bioink hydrogel and 72 then cell were exposed to RA after 1 day of culture. RA was added to construct for 48 hours by changing it every day. Finally, similar phases of differentiation were successfully obtained in the construct as they were observed in 2D: the cells first aggregated in the form of colonies, and after a few days they start to produce elongation of axons (Figure 5.1b). 73 Figure 5.1. Optimizing the differentiation protocol for neuronal differentiation in (a) 2D and (b) 3D. RA in defined medium (MEM/F12, 1% ITS, P/S) was added for 48 hours. At this stage, neural stem cells dominate the culture in dense planar sheets of culture (stage 1). After RA removal, dense aggregated colonies start to form in the culture (stage 2, a(i), b(i)). In 2D, these aggregates are loose and gradually detach from the surface, forming floating neural spheroids. In 3D, the spheroids form inside the printed fibers through migration and aggregation of cells (stage 3, a(ii), b(ii)). After 7-8 days of in vitro culture, the aggregated spheroids settle in the 2D culture and attach to the surface followed by neurite formation (stage 4, a(iii), b(iii)). Starting day 10 of differentiation, the neurites and neural networks are formed in the 2D culture. In the 3D culture, the neurites and neural networks are formed across and around the neuro-spheroids. Neurons are further matured by axonal development and high expression of specific biomarkers such as beta III tubulin (stage 5, a(iv), b(iv)). If FBS-containing culture media is supplied to the culture, the neurites fasciculate and the population will gradually get populated by neuroglia (stage 6, a(v), b(v)). To avoid this condition, neuron maturation media such as neurobasal medium or defined medium can be used. Scale bar in all images is 200 μm. 74 5.3. Differentiation study in 3D printed constructs After printing, the construct was incubated for 14 days of culture in the incubator. In order to study and the morphology of neuron cells in hydrogel and determine the presence of the neuron cells in the hydrogel fibers, the neuron cells were stained with beta tubulin III at day 7. As shown in Figure 5.2, the neuron cells printed in a construct expressed with beta tubulin III which is a specific marker for the neuron cell. This shows that NE-4Cs successfully differentiated to the neuron cells in the printed hydrogel. It is also observed from Figure 5.2 that the neurons were well defined, stayed in the fiber structure, formed a uniform channel fiber, and did not migrate to the supporting hydrogel. The neuron cells were stained after 14 day of culture. As shown in Figure 5.3, the neuron cells were stained with actin/DAPI and beta tubulin. The actin/DAPI shows the morphology of neuron cells after 14 days and beta tubulin shows that NE-4C differentiated to neuron cells. However, It was not possible to see the neurite elongation in the construct. All the cells expressed beta tubulin III marker which shows the presence of neuron cells and confirms that neural stem cells differentiated to neurons but not good spreading morphology of neuron. Three factors are considered that might have effect on the morphology of cells in this bioink material. First, the bioink material had to be optimized to be capable of neuron cell growth support. Second, the number of encapsulated cells had to be increased to have enough number of cells at differentiation stage. As the efficiency of differentiation is highly dependent on the cell density after RA induction [110], it is critical to optimize the cell density at the encapsulation stage to maintain a viable concentration of neural stem cells after printing. The third factor is proliferation of cells in the structure in which we culture the cells for two days before adding RA. 75 Figure 5.2. Beta tubulin/DAPI staining of neuron cells after 7 days of culture. a) DAPI channel, b) Beta tubulin channel, c) merged channel. Figure 5.3. a) F-Actin/DAPI staining of the neuron cells after 14 days of culture. b) Beta tubulin staining of the neuron cells after 14 days of culture. 76 For the optimizing the number of cells 25 million cells/ml was encapsulated in GGA bioink. As the efficiency of differentiation is highly dependent on the cell density after RA induction [110], it is critical to optimize the cell density at the encapsulation stage to maintain a viable concentration of neural stem cells after printing. Embedding stem cells at a concentration of 25 million/ml fulfilled the requirement for GGA bioinks. Prior to adding RA to the printed constructs, the 3D printed construct was allowed two days of in vitro culture to recover from possible loss of cell viability as well as allowing neural stem cells to proliferate further inside the channels. Increased metabolic activity of the neural ink confirmed that the viability reduction due to factors such as extrusion-induced shear stress can be fully recovered after a few days of culture in vitro. Moreover, a significant increase was observed in the metabolic activity of the neural cell-laden constructs from day 7 to day 14 (Figure 5.4), suggesting that the ink is capable of neural cell growth support. Thus, after 7 days of incubation, neural cells showed good spreading morphology in the printed nerve fibers (Figure 5.5a). By keeping the neural stem cells in culture until day 14, confluent fibers could be observed with elongated morphologies suggesting that the ECM properties of the bioink was suitable for cell growth and elongation (Figure 5.5b). 77 Figure 5.4 Metabolic activity of neural cell-laden bioink over the 14-day culture period Figure 5.5. Fluorescent images of neural cell-laden fibers labeled by F-actin and DAPI after day 7. For optimizing the bioink material: first different concentration of alginate (0.25 wt%, 0.5wt%, and 0.1wt %) were tested. The differentiation process was performed on all the different condition. As shown in Figure 5.6 significant difference was observed between the three. Alginate 0.5% was chosen for its best printability. Inks containing 78 Alginate with a concentration above 1% showed compromised printability and were not selected. Figure 5.6. Studying the effect of Alginate concentration in the bioink on neuron specific marker expression. Three ink conditions of alginate were printed and imaged after 10 days of differentiation. So significant difference was observed between the three. Alginate 0.5% was chosen for its best printability. Inks containing Alginate with a concentration above 1% showed compromised printability and were not selected. To better improve the bioink, further study is required. At the initial stages of differentiation (day 3-5), neural stem cells inside the channels transform into loose aggregated spheroids, a condition also observed in 2D. Upon proper surface functionalization (e.g. Poly-D-Lysine) in 2D cultures, such structures begin to settle in the bottom of the wellplate and spread neurites (Figure 5.1b(iv)). It was shown that addition of laminin to 3D gels can promote neuronal differentiation and neurite formation [111]. Therefore, laminin was added to promote the connections between neurospheroids. Reaching day 10 of differentiation, neuron structures start to form neuronal networks with 79 bundled axonal bodies. In 3D gels however, this process is further prolonged due to the different stiffness of the hydrogel ECM. In fact, the extension of the neurites inside a 3D matrix is critically dependent on the mechanical properties of the surrounding ECM [2]. Nonetheless, the 3D printed neural cells showed increased neuron-specific marker expression from day 7 onward (Figure 5.7). Following day 14, spheroids covered the neural fibers (Figure 5.8), and in some sections of the printed fibers, the neurons started to gradually migrate out the spheroid morphology inside the channel (Figure 5.8). It was hypothesized that the spreading of neurons inside the bioinks is enhanced through gradual degradation of bioink in culture; during the in vitro culture, the K+ ions in the culture medium gradually replaced the chelated Ca2+ in the egg-box structure of alginate chains resulting in degradation of the crosslinked alginate chains.[112] 80 Figure 5.7. Neuron-specific biomarker staining at different stages of differentiation. (a) At stage 1 and before differentiation induction, neural stem cells dominate the population and no βIII tubulin expression is observed in the 3D neuro-spheroids. (b) Upon differentiation induction, some cells (10%) start to show neuron biomarker expression. As the differentiation process is progressed, the density of neurons in the culture is increased to (c) 40% and (d) 80% defining stage 5 of neuronal differentiation. After this period, neuron formation is continued, and neurites and axonal bodies are gradually formed while some cells also differentiate to astrocytes in presence of FBS containing medium (e). 81 Figure 5.8. (a) phase contrast image of the neuron-laden fiber 20 days after neural differentiation induction, showing neuron spreading (b) βIII tubulin and DAPI staining of the aligned printed fibers (boundaries shown in red) after 14 days of culture on day 10 of differentiation. (c,d) an internal network of neurons in the printed nerve fibers expressing positive βIII tubulin and negative GFAP which shows successful differentiation of neural stem cells into neurons and not neuroglia. 5.4. Summary In this chapter, differentiation of neural stem cell into the neurons was studied in single culture setting. The study was performed and optimized in 2D and then extended into the 3D. Overall, differentiation of stem cells was achieved in fiber-like 3D format after 14 days of culture. 82 Chapter 6 Co-culture 6.1. Co-culture of Neuron stem cell and Astrocytes The final aim of this thesis is to print neural stem cells fiber into supporting bath containing astrocytes in order to co-culture the neuron and astrocyte in a printed structure. A main consideration in co-culture tissue models is to optimize the common media between the different cell types. Effective differentiation of neural stem cells into neuronal subtypes requires a serum-free conditioned media while the glial culture rely on serum-containing media for proliferation. Studies which target neuron-glia co-culture models have used different methods for media supplement. For instance, Aregueta-Robles et al mediated the co-culture media by using a half-half composition of complete media specific to each cell type [113]. Other studies showed that astrocytes can survive and be cultured in the neuron-maintaining medium[114, 115]. Therefore, three different media known to enhance neuronal differentiation were used and tested for astrocytes growth. These three different media were; Neurobasal medium supplemented with B27 (NBM), Defined Medium (DM, containing MEM/F12+1%ITS+1%P/S), and neural stem cell media (NS, containing MEM+10% FBS+1% P/S). For the control media, DMEM (DMEM, 10% FBS, 1% P/S) a known media for astrocytes were used. 83 6.2. Stage 1: optimization of the differentiating media for astrocyte growth in 2D A study was performed to optimize the common co-culture media composition for both neuron differentiation and astrocyte growth (Figure 6.1). Astrocytes were cultured in various neuron-supportive media including Neurobasal medium supplemented with B27 (NBM), defined medium (DM, containing MEM/F12+1%ITS+1%P/S) as recommended by previous research [116]. For control groups, we chose astrocyte media (DMEM, 10% FBS) and neural stem cell media (MEM, 10% FBS). After 7 days of culture, astrocytes showed good survival in the DM medium as opposed to the NBM group. High expression of GFAP marker confirmed the survival and preservation of astrocyte phenotype in the DM media. As a result, DM was chosen as the common media for maintaining all co-culture constructs. 6.3. Stage 2: optimization of the media in 2D co-culture After optimizing the co-culture media for neural stem cells and astrocytes. Neural stem cells and astrocytes were seeded on PDL coated well plate with the concentration of 2 million cell/ml. RA in defined media (MEM/F12+1%ITS+1%P/S) which optimized for co-culture was induced for 48 hours. Neural stem cells started to differentiate following the different stages as described in Chapter 5. Neuron cells and astrocytes were stained with their specific marker (beta tubulin and GFAP respectively) at day 7 (see Figure 6.2). As shown in Figure 6.2, neural stem cells were differentiated to neuron and are at the stage 4 of differentiation (as described in section 5). The aggregated spheroids of neurons settled 84 in the 2D culture and attached to the surface followed by neurite formation. The astrocytes also proliferated, connected to each other and made network (Figure 6.2). 85 Figure 6.1. Optimization of co-culture media for brain-like co-culture constructs. Astrocytes were cultured in three different media known to enhance neuronal differentiation. Neurobasal medium supplemented with B27 (NBM), Defined Medium (DM, containing MEM/F12+1%ITS+1%P/S), and neural stem cell media (NS, containing MEM+10% FBS+1% P/S) were tested. DMEM astrocyte media (DMEM, 10% FBS, 1% P/S) was selected as the control culture. Astrocytes cultured in NBM did not survive the one-week culture. Cells cultured with DM, however, showed good proliferation. As such, DM was selected as the common media for co culture of neuron- and astrocyte- laden constructs. GFAP/DAPI staining (right column) confirmed this observation. Scale bar in all images is 200 μm. 86 Figure 6.2. Co-culture fluorescent images of neuron and astrocyte cells labeled by BIII tubulin and GFAP individually on 2D after 14 days of incubation 6.4. Stage 3: optimization of the media in 3D co-culture In order to study the feasibility of co-culturing the neuron and astrocytes in optimized biomaterials as well as optimizing the cell density and culture time for neuron-astrocyte co-culture, astrocytes and neural stem cells with different cell density were first encapsulated in a thin layer of GelMA/gelatin/ alginate/laminin hydrogel. The thickness of hydrogel was 2 mm. RA in defined media (MEM/F12+1%ITS+1%P/S) was induced for 48 hours and defined media (MEM/F12+1%ITS+1%P/S) was used for further culture. Neuron cells and astrocytes were stained with their specific marker (beta tubulin and GFAP respectively) at day 7 (see Figure 6.3). Astrocytes were proliferated and started making network and connection as early as day 7 (Figure 6.3). However, the neuron cells only express beta tubulin marker at day 7 and the morphology of them is not like a mature neuron cells. It is obvious that, the longer culture time for 3D compared to 2D culture is 87 needed for neuron cells to differentiate, show their morphology and started interconnecting. The results shows that the cell density of neural cell optimized for 25 million/ml and astrocytes for 8 million/ml and culture tine more than 7 days is required for neural stem cells to differentiate into neuron in 3D structure. Figure 6.3 Co-culture fluorescent images of neuron and astrocyte cells labeled by BIII tubulin and GFAP individually in a thin layer of hydrogel 14 days of incubation. 6.5. 3D bioprinted brain model co-culture At the final stage, 25 million cells /ml were encapsulated in GelMA/ gelatin/ alginate/ laminin bioink and printed in the GelMA supporting bath containing astrocytes with the density of 8 million/ ml. The printed co-culture construct kept culturing for two days to allow neuron cells to proliferate more. Then RA in defined media was added to the 88 construct for 48 hours to induce neural stem cell differentiation. The co-culture printed construct were kept for 30 days culturing and every 3 days the construct were observed under microscope. The neural stem cells were differentiated into neuron cell following different stages (see Figure. 5.7) and astrocytes started growing and making network. As shown in Figure 6.4, similar to the result of neural cell-laden co-culture, on day 7, a well elongated neural cell proliferation was observed while the astrocytes showed a less spindle morphology around the printed fibers. It is hypnotized that the difference between the physically crosslinks alginate chains and chemically photo-crosslinked GelMA chains have created two rates of extrusion in the cells. Figure 6.4. Co-culture fluorescent images of neuron and astrocyte cells labeled by BIII tubulin and GFAP individually in the brain-like structures after 14 days of incubation 89 6.6. Summary In this chapter, the co-culture media was optimized for neural differentiation in co-culture settings. Neutral stem cells was differentiated into the neuron in the presence of astrocytes. The differentiation was achieved in the printed fiber like structures proposed in the fabrication chapter. 90 Chapter 7 Conclusion & suggestions for future work This chapter covers a summary of the thesis, the major contributions to the field and possible future directions for expand the current study. 7.1. Summary Here, a novel brain model was developed based an embedded bioprinting that mimics the brain’s fiber-like structure and includes two main cell-types in the CNS. The fabrication strategy was then carried out, optimized and characterized in detail to create robust and high throughput biofabrication methodology. Then several physical and biological characterization studies were carried out to optimize the proposed biomaterials for printability, robustness and gelation. The materials were also optimized to achieve high cell viability and suitable cellular activities in printed cell-laden constructs. Finally differentiation studies were carried out to discover a methodology for biofabricating the neural model based on neural stem cells. The printed models were characterized through immunostaing, physical and biological quantifications. 7.2. Contribution to the field I developed a biofabrication strategy based on a 3D embedded printing technique to devise a 3D brain model which closely mimics the native brain structure. The following lists describe the novelty and potential features of such a model:  This is a high-throughput model which could be used for drug testing. 91  This is the first model which mimics the geometric of brain. The fiber structure of neurons is surrounded by astrocytes in 3D. The model could provide an insight into brain development and functionality.  This is the first 3D in vitro model containing astrocytes and neurons cells positioned in 3D structure. Therefore, this model could recapitulate the neuron-astrocyte interaction to study and control the levels of glutamate in the brain. As such, the engineered brain-like co-culture constructs can be potentially used as a reliable and highly reproducible in vitro platform for brain disease modeling.  The model contains neural stem cells which are differentiated to neurons, making the model cost effective for CNS stem cell therapy. 7.3. Future work The model proposed in this thesis was based on neural stem cells differentiation into neurons in the presence of astrocytes. This, however, proved to be difficult to achieve, as optimizing cell culture media in co-culture of astrocytes and neural stem cells requires a lot of trails and further studies. Therefore, here as a future direction, I propose to use isolated neural cells from mice instead of neural stem cells. Note that, a single isolation of neurons from mice can be used to print multiple devices suitable to study the brain basic functions and also they can serve as a disease model. Meanwhile optimizing the cell media is much easier in this case, as the differentiation is not required. Another proposal is to use neuroblastoma cell types in the model instead of neural stem cell to model for brain tumors that mimics the 3D geometrical complexity of the 92 brain. Here, the model can be used to test different type of drugs aimed to regulate neural functions. Final suggestion is to add hollow vessel into the model to create blood vessel in the thick hydrogel construct which enhances the model and functionality of printed model. Meanwhile it can pave the way to add endothelial cells to model blood vessels in the brain to create more accurate model and move toward modeling blood-brain-barrier (BBB). The addition of hollow channels also facilitate fabrication of thick constructs and can provide a mean to introduce flow in the model to create an organ-on-a-chip system. 93 Appendix A Engineering biocompatible hydrogel fibers with tunable mechanical properties for neural tissue engineering Tissue engineering holds great promise for treatment of neurodegenerative disease. In this study, a microfluidic system was used to develop core-shell hydrogel fibers. The hybrid hydrogel was formed from GelMA mixed with gelatin containing cells in core and alginate in shell. The composition of the core hydrogel was optimized to support cellular growth and differentiation, yet allow feasible fabrication of cell-laden fibers. The engineered fibers were remarkably biocompatible and enabled the formation of highly organized cellular morphology [12, 15, 117-136]. A.1. Introduction The human central nervous system (CNS) has a limited capacity to counteract the loss or dysfunction of neurons and axonal pathways that accompany conditions such as traumatic brain injury (TBI), stroke, spinal cord injury (SCI), and neurodegenerative disease [64, 137-140] . Neurogenesis in the CNS is restricted to a limited number of areas in the brain, hampering the restoration of lost neurons [141, 142]. Additionally, regeneration of lost axonal pathways in the CNS is insufficient due to the lack of directed guidance. Commonly used/ conventional regenerative strategies including stem cell transplantation/engraftment or growth factor delivery have shown limited success in creating functional tissues. The key challenges in such therapies have been low cell engraftment and lack of growth factors localization, which have resulted in[143]. Recent 94 advancements in neuronal tissue engineering have resulted in some improvement to scaffolds for neuronal cell growth, however, the success of these acellular therapies depends on the presence and recruitment of healthy cells to the injury site. Tissue engineering strategies based on engineering cell-laden constructs have shown great promise for forming functional tissues or restoring the function of damaged tissues [144, 145], however, most scaffolds/constructs (have the following weaknesses). Researchers have employed electrospinning, molding, photolithography, and hot embossing to generate 2D patterns [146-151]. However, these techniques fail to mimic the mechanical properties and architecture of brain tissues. Some recently developed fiber-based technologies such as biotextiles and 3D bioprinting have enabled the fabrication of tissue constructs with anisotropic properties and architectures [152-155]. The key component of fiber-based tissue engineering strategies the fabrication of cell-laden fiber scaffolds that can be assembled to form a 3D structure [156, 157]. Thus, creating functional microfibers carrying neuron cells followed by their assembly could result in the formation of functional 3D neuronal tissue. Here, we formed alginate-based hydrogel, which can easy be spun into fibers. The alginate-based hydrogel was used to fabricate core-shell fibers in which the core was formed from the biocompatible hybrid hydrogel comprised of gelatin methacryloyl (GelMA)/gelatin and hydrogel that contained cells, while the shell was made from alginate. The fibers were mechanically robust and can be assembled using textile processes into 3D constructs. Furthermore, the fibers were used for culturing neuroblastoma cells and it was shown that they could support the growth and reorganization of neuroblastoma cells with an aligned structure. 95 A.2. Material and Methods Materials Sodium alginate, calcium chloride (CaCl2), and DAPI were purchased from Sigma Aldrich (St. Louis, MO, USA). 2-hydroxy-1-(4-(hydroxyethoxy) phenyl)-2-methyl-1- propanone (Irgacure 2959, CIBA Chemicals) was used as photoinitiator (PI). Dulbecco’s modified Eagle medium (DMEM), fetal bovine serum (FBS), 0.05% trypsin-EDTA (1X), Live/Dead Assay kit, PrestoBlue® Cell Viability Assay, and antibiotics (Penicillin/Streptomycin) were purchased from Invitrogen (Carlsbad, CA, USA). GelMA was prepared according to our established protocol [158, 159]. GelMA/Gelatin hybrid hydrogel preparation Hybrid hydrogel was prepared by mixing the same volume of 2% (w/v) gelatin (in HBSS), 3% (w/v) GelMA (in HBSS) and finally adding 1% (w/v) PI (in HBSS). Subsequently the hybrid solution was mixed for 10-20 secs with the vortex mixer in low speed. Fabrication of hydrogel fibers To fabricate the hydrogel fibers, a co-flow device was utilized by assembling two needles of different sizes together. The needles were chosen in a way that the outer dimeter of one is less than the inner diameter of the other, where there would be a space between the needles for liquid to flow after assembling them together. Here we used needles of 18G and 20G sizes to make the device. The head of the needles were trimmed and flattened using sandpaper. An inlet (focusing flow inlet) was made in the plastic chamber of the bigger needle. The needles were then assembled together and glued. A transparent 96 chemical resistant clear PVC tubing was then connected to the outlet and placed in a 2% calcium chloride solution in water. As it is illustrated in Figure A.1, hydrogel fibers were fabricated using inner phase of 3% GelMA mixed with 2% gelatin (different concentration were made), and the focusing flow phase of 2% sodium alginate. The two phases were injected into the inlets of the microfluidic device. The flow rates of the solutions were finely controlled by syringe pumps (PHD 2000, Harvard Apparatus) to create fibers with different sizes. The hydrogel fibers were collected in a Petri dish containing 2% calcium chloride in water to fully crosslink the hydrogel. The fibers were then exposed to UV light (OmniCure s200, USA) for 30 secs to crosslink the GelMA hydrogel in the fibers. The UV intensity was set to 850 mW and the distance between the tip of the fiber optic and the Petri dish was set to 8 cm (see Figure A.1). Cell culture Neuroblastoma (SH-SY5Y) (ATCC, USA) were cultured in a 1:1 mixture of Dulbecco’s Modified Eagle Medium (DMEM, Sigma) and F12 medium supplemented with 10% (v/v) fetal bovine serum (FBS, Sigma) and 1% (v/v) penicillin-streptomycin (Gibco, USA). Cells were maintained at 37ºC in a humidified 5% CO2 atmosphere until 80-90% confluence was reached. Cell passages 6-8 were used during experiments. SH-SY5Y cells encapsulated in pre-described hybrid hydrogel using the following steps; after detaching cells (by trypsin), they were resuspended in cell culture media with the density of 15 M/mL. Cells solution was added to GelMA and gelatin hydrogel solution. Subsequently GelMA/glatin containing cell solution was mixed with PI solutions. 97 Viability Assay (Live/Dead) To evaluate the in vitro biocompatibility of the obtained hydrogel fibers, SH-SY5Y encapsulated in the shell of hybrid hydrogel fibers were cut into smaller segments (1 cm) and placed in a 12-well plate. Viability test was performed by incubating cell-laden samples for 15 minutes at 37°C with a mixture of 2 µl/ml ethidium homodimer-1 (EthD-1, red, dead cells, Invitrogen) and 0.5 µl/ml calcein AM (live cells, green, Invitrogen). Fluorescence Microscopy was performed on a Zeiss Observer.D1 microscope using an X-Cite Series 120Q fluorescence source. Actin/DAPI staining Samples were fixed in 4% paraformaldehyde (Electron Microscopy Sciences), washed in DPBS and subsequently stained with phalloidin (life Technologies, labeled with Alexa F 594) and DAPI as described in the manufacturer’s manual. Microscopy was performed as described above. A.3. Results and discussion A.3.1. Fabrication of Hybrid Hydrogel Fiber The schematic of the microfluidic setup used to fabricate hydrogel fibers is presented in Figure A.1A. The system consisted of two syringe pumps, a co-flowing microfluidic chip with two inlets and one outlet, and a calcium chloride (CaCl2) bath. The microfluidic system was made by two concentric metallic needles together to create core and sheath streams of different fluids at the outlet of the device. The outlet of the microfluidic device was connected to the bath with a capillary tube 150 µm in inner diameter (ID). To fabricate the composite cell-laden hydrogel fibers, a solution of sodium alginate and a solution of GelMA/gelatin containing the neuroblastoma cells (see material 98 and method section for detail of the solutions) were pumped through the microfluidic device as the shell and core fluids, respectively (Figure A.1.C). GelMA possesses cell binding sites [158]. However, wet spinning of GelMA is challenging [157, 160]. Tamayol et al. proposed the use of alginate to form a template and physically entrap GelMA monomers in the fiber template prior to their photocrosslinking into a polymeric network [160]. A similar approach is used here to create the shell of the composite fibers. Although alginate lacks cell binding moieties, it has been shown that the use of suitable concentration of alginate does not affect cellular function. Thus, once the solutions extruded into the CaCl2 bath, calcium ions replaced the sodium and created calcium alginate hydrogel [161]. Finally, the collected fibers were exposed to UV light to photocrosslink the GelMA hydrogel. The result of the fabrication process is a robust hydrogel containing cells fiber and a hydrogel shell (Figure A.1B). The diameter of the core and the thickness of the shell could be tuned simply by controlling the ratio of the flow rates. The fabricated fibers were biocompatible and the encapsulated cells showed elongated morphology. (see Section 2.4 for more details). 99 Figure A.1. Conceptual view of the high throughput fabrication process of core-shell hydrogel fiber and the resulted construct. (A) The setup was consisted of a microfluidic device with the ability to control the flow rate of core and shell solution, separately. Extruding the solutions through co-axial microfluidic device can induce shear force within the fluid and elongate polymer (i and ii). The structure then was injected to CaCl2 bath where alginate was crosslinked to form the fiber’s matrix. Crosslinking alginate trapped the hydrogel in the core (GelMA/gelatin) and formed a template network in the shell (iii). (B) Finally, the samples were exposed to UV-irradiation for crosslinking GelMA. Representative micrographs showing produced fiber. Scale bar is 500μm. (C) schematic of neural fiber generation, alginate acts as a template preventing neurons from diagonal expansion and force them to within the GelMA/gelatin core and along the fiber, which leads to interconnection between neurons. A.3.2. Physical properties The core-shell fibers are fabricated by microfluidic set up (Figure A.2A). The two axial nozzle consisted of two different needles which concentrically assembled together to make the different channels for the core and shell flows. GelMA/gelatin was pumped to 100 the core channel and alginate was pumped to shell channel through the syringe pumps and the fabricated fibers introduced to CaCl2 (Figure A.2A). The constructs in Figure 2 are fabricated by the use of 3%/2% (w/v) of GelMA/gelatin ink for the core and 2% (w/v) alginate ink for the shell. The fabricated structures maintained good resolution and the core and shell regions remained un-mixed (see Figures A.2B). The flow rates of different streams were changed to investigate their effect on the dimensions of the fabricated fibers. By increasing the flow rate ratio between the inner streams to the outer one, the diameter of printed hydrogel fiber was increased (see Figure A. 2C). The mechanical properties of the fabricated fibers could be tailored by changing the properties of the used pre-polymers. Mechanical properties of the fibers could also be tailored by changing the alginate concentration. Higher alginate concentration yielded larger young modulus (Figure A.2D). The young modulus of core hydrogel material consisted of GelMA/ alginate were measured for different concentration of GelMA and 1% gelatin. The results show that by increasing the concentration of GelMA, the young modulus is increased. (Figure A.2E) The young modulus of the 2.5% GelMA and 1 % gelatin is about 0.6 kpa which is around the stiffness of the native brain tissue. Therefore, this concentration is picked for the core of hydrogel. 101 Figure A.2. Characterization of hydrogel fibers. (A) fabrication setup of core-shell hydrogel fibers (B) Geometrical features of core-shell hydrogel fibers as compared to sold hydrogel fiber (C) The diameter can easily be controlled by setting the proper inflow rate of sheath and core fluids (D) Mechanical properties different alginate concentration (E) Mechanical properties of GelMA/gelatin hydrogel with variation of GelMA concentration, gelatin was kept constant as 1% for all (*: P<0.05) A.3.3. Cytotoxicity evaluation To evaluate the ability of the materials as well as the fabrication method to provide suitable environment for cells growth and proliferation, three different tests were carried out using Neuroblastoma cell line (Figure A.3). Cells were encapsulated in GelMA/ gelatin hydrogel and fiber structure were formed using co-flow device. The cells were stayed in the core since the alginate hydrogel prevents cell to migrate into that. The Figure A.3A shows the phase contrast images of cells encapsulated in hydrogel fiber at day 0, 1 and 5. Viability were assessed using live/dead assays (Figure A.3B). Live/dead assay showed 102 more that 85% of the cells were live (appearing as green) after one and three days of culture. The results suggested that the fabricated fibers do not induce immediate toxicity and also the presence of alginate with the tested concentrations did not reduce cellular viability. In addition, F-actin/DAPI staining was performed for the cells encapsulated in the hydrogel core (Figure A. 3C) to evaluate the cells morphology in 3D structure of hybrid hydrogel. Findings demonstrated a good proliferation in highly organized patterns along the fiber axis. The results further confirmed the biocompatibility of the engineered hydrogel. Figure A.3. Viability and proliferation of neuroblastoma encapsulated in fiber. (A) Phase-contrast images of neuroblastoma encapsulated in hydrogel fiber in day 0, day 1, and day 5, scale bar 100 um. (B) Live/ dead staining of neuroblastoma at day 1 and day 3, scale bar 100 um. (C) F-actin/DAPI staining at different days of culture, day 1, day 5, day7, and day14, scale bar 50 um. 103 Cellular alignment is critical in proper function of brain tissue. We observed that the hybrid hydrogel system significantly promoted cellular alignment along the fibers. Figure A.4 shows qualitative and quantitative results of the cellular orientation in different days. This oriented spreading helps the final goal to make neural fiber mimicking the real construction of the brain tissues. Fluorescent staining of neuroblastoma F-actin filaments of Figure A.4A indicates highly aligned cell with respect to the fiber orientation while these patterns maintain during further spreading and proliferation. The orientation of F-actin filament was quantified and measured with respect to the direction of the channel. The results showed that after 7 days actin filaments were oriented in the direction of the channel with the standard deviation of less than 10 deg. The orientation was improved on Day 14 with the standard deviation of less than 6 deg. (Figure A.4B). The cell culture density within the 3D tubular structure was estimated on Days 7 and 14 per unit volume of the channel to compare the seeding density and those of the cells on the following days. The results showed that the seeding density was 15 million/ml; whereas the density of cells after 7 and 14 days of culture was increased to 27 and 50 million/ml, respectively (Figure A.4C). The high density of the cells can be atributed to the highly-oriented actin filaments observed on Days 7 and 14, as the higher density of the cells causes elongation of the cells in the direction of the channel. We speculate that cellular alignment might be due to three phenomena: 1) the morphology of the fibers and the stiffness of the core fiber which applied mechanical stimulation to cells to align them in concentric patterns; 2) the alignment of polymeric sheets due to the flow induced shear stress during the extrusion process; 3) the alignment of unmixed GelMA domains. GelMA and alginate do not mix homogeneously due to their dissimilar physical properties such as density, and thermal 104 sensitive viscosity; since alginate hydrogel does not provide a cell friendly matrix, cells align toward the elongated GelMA networks (see Figure 1). It has been shown that the size of hydrogel constructs play a key role in the morphological growth of cells [162, 163]. In smaller GelMA features cells can sense the boundaries and align themselves with them. In current fibers, the core material is much stiffer than the sheath cell-carrying niche and the small thickness of the cell-laden niche resulted in alignment of cells along fibers axis. Figure A.4 Qualitative and quantitative analysis of cellular orientation in different days of the culture for the encapsulated cells in fiber. (A) F-actin/DAPI staining after 7 and 14 days, respectively from left to right. (B) Deviation of neuroblastoma from the channel direction for Days 7 and 14. (C) Cell seeding density and culture density after Days 7 and 14. A.4. Conclusions We engineered cell-laden hydrogel fibers by developing hydrogel composites with tunable mechanical properties. A microfluidic system was used to fabricate core-shell fibers with a cell-laden core and shell from a biocompatible cell-laden hybrid hydrogel of GelMA/gelatin. The hydrogel fibers were biocompatible and supported the growth of encapsulated neuroblastoma cells. The fibers also positively directed cellular morphology and enabled the generation of highly organized cell networks. The developed materials 105 are mechanically robust and can be used in 3D bioprinters to fabricate 3D constructs. They can be easily assembled into 3D structures with clinically relevant dimensions using a combination of 3D printing and textile processes. These engineered fibers can be potentially used for tissue engineering in vitro models for studying neuronal development or drug testing. 106 Appendix B Ferritin Nano-cage Conjugated Hybrid Hydrogel for Sustained Drug Delivery and Tissue Engineering Applications Hydrogels have been extensively used in tissue engineering scaffolds, drug delivery systems and wound healing applications due to their unique compositional and structural similarities to the natural extracellular matrix (ECM), high water content and biocompatibility. Despite enormous advances in the application of hydrogels in various biomedical applications, poor mechanical properties and lack of control on release behavior of biomolecules act as major barriers for widespread clinical application of hydrogel[164-167]. Here, a ferritin incorporated gelatin methacryloyl (GelMA), nanocomposite hydrogel with improved mechanical properties, cell responsiveness, and sustained-release behavior for tissue engineering and drug delivery applications was presented. Ferritin and its empty- core equivalent apoferritin, were chemically modified to present methacryloyl groups, ferritin methacryloyl (FerMA) and apoferritin methacryloyl (ApoMA) respectively[168]. FerMA and ApoMA were covalently conjugated to gelatin methacryloyl (GelMA) hydrogel via UV crosslinking. The developed hybrid nanocomposite hydrogels offered better ability to tune the mechanical properties compared to those prepared by direct dispersion of ferritin and apoferritin into GelMA hydrogel, without affecting their porosity or cell growth. The NIH-3T3 cells showed improved spreading on the developed hybrid nanocomposite GelMA hydrogel. The ability of the nano-cage to deliver small biomolecule compounds was confirmed by performing 107 cumulative release test on apoferritin/ApoMA-GelMA hydrogel with FITC core. Thus, the apoferritin/ApoMA-GelMA hydrogels have emerged as an excellent and promising nanocage-hydrogel for sustained drug delivery, and tissue engineering applications, thanks to their unique architecture, surface properties and high biocompatibility. B.1. Introduction Hydrogels have been widely used in numerous biomedical applications including drug delivery, wound healing, biomaterials and tissue engineering applications. They are a unique class of biomaterial, of hydrophilic nature and polymeric three-dimensional (3D) networks that can absorb large amount of water and permit the diffusion and attachment of molecules and cells[169]. Due to their high biocompatibility and the similarity of physical properties to natural extracellular matrix (ECM), hydrogels have recently drawn a lot of attention for applications in the field of biomedicine such as tissue regeneration[170], organ-on-a-chip[171], biosensors[172] and drug delivery areas[173]. However, there are still some clinical and pharmacological challenges such as low tensile strength and uncontrollability of drug delivery that are needed to be addressed for wide spread biomedical applications of hydrogels [174-179]. Several approaches have been used to tune the physical and biological properties of hydrogels to match the criteria required for their specific tissue engineering application and drug delivery, such as chemical modifications[180] and nanoparticle incorporation[170]. In particular, a range of nanoparticles such as carbon-based nanomaterials[181] (carbon nanotubes (CNTs), graphene), metal/metal-oxide nanoparticles[182] (gold, silver, copper, platinum, iron oxides, and titanium oxides), inorganic/ceramic nanoparticles[183] (silica, clay, silicates, and calcium phosphate), and organic nanoparticles[184] (polymer 108 nanoparticles, micelles, polymersomes, dendrimers, and liposomes), have been integrated within the hydrogel networks to gain nanocomposite hydrogels with tailored functionality and superior biological and physicochemical properties. For examples, silica nanoparticle incorporated hydrogels showed remarkable improvements in mechanical stiffness, bioactivity, and tissue stickiness compared with unmodified hydrogels [185, 186]. In addition, silver and gold nanoparticles and CNT were also incorporated into hydrogel structure to enhance mechanical properties, raise the electrical conductivity or add/promote antibacterial activities [170, 181, 182, 187] . Among the various types of nanoparticles that can improve the mechanical and biological properties of hydrogels along with additional functions such as delivery of small compounds and bioactive molecules, ferritin nano-cage is an excellent candidate with various desirable features[188]. Ferritin, present naturally in almost all living organisms where it functions as a major iron carrier, is a globular protein complex consisting of 24 protein subunits which self-assemble into a spherical cage-like structure[189]. Ferritin has a protein shell of about 2 nm surrounding an iron core of about 8 nm which can store about 4500 iron atoms and release them in a controlled fashion. Its iron free protein shell, apoferritin, is a pH-dependent hollow spherical structure which was stable in pH 4–9 aqueous environments[190]. Apoferritin reversibly disassembled when the pH becomes extremely acidic (pH 2–3) or basic (pH 10–12)[190]. When the pH returns to neutrality, ferritin monomers can associate in a shape memory fashion. Apoferritin is highly resistant to chemical and physical denaturants, including heating to high temperatures (>80 0C). Owing to their uniform nanosize, function stability, core-shell structure and unique surface properties, apoferritin has emerged as an excellent and promising protein-based nanocage. 109 Apoferritin has shown to be able to capture different drugs/nanoparticles and release them in specific condition (e.g. change of pH), such as MRI contrast agents, antibacterial silver nanoparticles, gold nanoparticles and various chemical catalysts[191]. Intrinsically, ferritin nanocages may guarantee a proper drug delivery and release by exploiting natural recognition of the Transferrin Receptor 1, which is overexpressed on tumor cells[192]. Moreover, ferritin can be either chemically or genetically modified to apply functionalities to their surface for further providing active tumor targeting. Study proved that the surface of ferritin can be modified with RGD-peptides to bind to a specific tumor, efficiently releasing drugs[193]. Furthermore, ferritin can form networks with polymer chains through the carboxylic acid and amino groups of shells[194, 195]. Previous study has demonstrated that fibers of PVA hydrogel decorated with ferritin had improved elastic modulus and mechanical strength, 230% and 170% respectively[194]. In this case, it was proposed that the peptide shell interacts as a spring, passing the force to the stable core and therefore enhances the mechanical properties. Ferritin has also been demonstrated to be used as an interfacial adhesion promoter. In this study, ferritin particles were covalently attached to carbon-nanotubes[196]. These MWCNTs functionalized with ferritin were readily incorporated in poly (vinyl alcohol) (PVA) hydrogel, solving the problem of weak interfacial adhesion between the nanotubes and the matrix. To create a ferritin reinforced hybrid hydrogel that can be used for biomedical application, gelatin methacryloyl (GelMA), a photocrosslinkable hydrogel, was utilized which has suitable biological properties and tunable physical characteristics and is one of the most popular investigated hydrogels[197, 198]. The cellular microenvironment plays a 110 crucial role in improving cell response and function of an engineered tissue. Due to the cell attaching and matrix metalloproteinase responsive peptide motifs, which allow cells to attach, proliferate and spread in GelMA-based scaffolds, GelMA hydrogels closely mimic essential properties of native ECM and are great candidates for tissue engineering applications[197]. Physical and mechanical properties of GelMA can be easily tuned by manipulating the type and ratio of its components. Consequently, ferritin incorporated GelMA hydrogels with drug delivery potential are expected to be a promising scaffolding material for 3D tissue engineering constructs. Here, a simple approach was reported to chemically modify the surface of ferritin/apoferritin with methacrylate (MA) groups to facilitate covalent bonding of ferritin/apoferritin with GelMA network while keeping its drug-delivery property. Firstly, the structure of the ferritin and apoferritin after methacrylation were verified using a transmission electron microscope (TEM). Mechanical properties, toughness, and elongation of the FerMA/ApoMA-incorporated hydrogel were compared with those made with unmodified ferritin/apoferritin to evaluate the effects of covalent conjugation. The biocompatibility of composited hydrogels was evaluated by measuring the viability and immunostaining of encapsulated fibroblasts. Finally, a FITC release strategy via apoferritin nanocage was tested as a proof of concept using apoferritin/ApoMA incorporated GelMA for drug delivery application. 111 B.2. Material and methods Materials Gelatin (Type A, 300 bloom from porcine skin), methacrylate anhydride (MA), 3-(trimethoxysilyl)-propyl methacrylate (TMSPMA), NaOH pellets, NaCl, ferritin and apoferritin both from equine spleen were purchased from Sigma-Aldrich (Sigma-Aldrich Corporation, St. Louis, Missouri, USA). The photoinitiator Irgacure 2959, 2-hydroxy-1-(4-(hydroxyethoxy)-phenyl)-2-methyl-1-propanone, is acquired from CIBA (Ciba Specialty Chemicals Incorporated, Basel, Switzerland). Dulbecco’s phosphate buffered saline (DPBS), Dublecco’s modified eagle medium (DMEM), fetal bovine serum (FBS) and penicillin-streptomycin were bought from Invitrogen (Invitrogen Corporation, Waltham, Massachusetts, USA). Dialysis membrane tubes, Slide-A-Lyzer Dialysis cassette, glass slides and coverslips were purchased from Thermo Fisher Scientific (Thermo Fisher Scientific, Waltham, Massachusetts, USA). DAPI (4’,6-Diamidino-2-Phenylindole), Alexa Fluor 488 Phalloidin, LIVE/DEAD viability/cytotoxicity kit and Fluoraldehyde o-Phthaldialdehyde (OPA) were acquired from Life Technologies, a Thermo Fisher Scientific brand (Thermo Fisher Scientific, Waltham, Massachusetts, USA). The UV source used (Omnicure S2000) is manufactured at EXFO Photonic Solutions (EXFO Photonic Solutions Incorporation, Quebec City, Canada). The device for mechanical testing is the 5942 Single Column Tabletop Model testing systems from Instron, an ITW company (Illinois Tool Works Incorporation, Glenview, Illinois, USA). 112 Synthesis of GelMA: Gelatin methacrylate (GelMA) GelMA was synthesized following a protocol based on previous works[197]. In short, 10 wt% gelatin from porcine skin (G1890, Sigma-Aldrich) was dissolved into Dulbecco's phosphate-buffered saline (DPBS) at 50 °C using a magnetic stirrer (240 rpm). 8 ml methacrylate anhydride (MA) (Sigma Aldrich) was added drop wise, in order to reach a final concentration of 8 wt% MA. After 2 h at 50 °C with constant stirring, the solution was diluted 1:1 with DPBS (previously warmed) and dialyzed against distilled (DI) water with dialysis membrane (Spectra/Por molecular porous membrane tubing, MWCO 12-14,000, Fisher Scientific) for 5 days at 40 °C, maintaining a constant stirring (500rpm) and changing DI water every 12 h. The dialyzed solution was stored at -80 °C and, according to experimental needs, freeze-dried for 5 days to prepare for use. Ferritin/ApoFerritin Methacrylation (FerMA/ApoMA): 1 ml Ferritin solution was diluted by the same amount of DPBS in a stirring (300 rpm) glass vial at cold room conditions (4 oC). 20 μl methacrylate anhydride was added dropwise and the solution was stirred overnight at the same speed for at least 12 hours. The FerMA was dialyzed with dialysis cassette (Slide- A-Lyzer, MWCO 12000, Fisher Scientific) against normal saline solution (9 g sodium-chloride dissolved in 1L distilled water) for 3 days. Apoferritin is delivered in a lower concentration (13 mg/ml) than Ferritin (52 mg/ml). Therefore, the Apoferritin solution did not need further dilution. The procedure for creating Apoferritin-methacrylate is the same as for creating FerMA. To 1 ml Apoferritin solution 20 μl methacrylate anhydride was added dropwise. The reaction and dialysis were done the same way. 113 Characterization of Methacrylation To qualify the methacrylation of primary amine groups on the protein surface of ferritin and apoferritin a commonly known fluoraldehyde assay was used. Fluoraldehyde o-Phthaldialdehyde (OPA) reacts with free amine groups (-NH2) giving a blue fluorescent product. The excitation and emission maximum are at 340 nm and 455 nm. The fluorescent signal can be measured, correlating directly with the number of free amine groups in the solution. Solutions with the same concentration of proteins (100 μg/ml) need to be prepared. The same amount of fluoraldehyde OPA reagent solution was added and in a quick manner the fluorescent spectrum was measured. The differences compared to a reference (DPBS, in which the proteins are diluted), were recorded and the degree of substitution calculated with the following equation: 𝑍 = (1 −𝐼𝑃𝑟𝑜𝑡𝑒𝑖𝑛−𝑚𝑒𝑡ℎ𝑎𝑐𝑟𝑦𝑙𝑎𝑡𝑒 − 𝐼𝑅𝑒𝑓𝑒𝑟𝑒𝑛𝑐𝑒𝐼𝑃𝑟𝑜𝑡𝑒𝑖𝑛 − 𝐼𝑅𝑒𝑓𝑒𝑟𝑒𝑛𝑐𝑒) × 100 In which Z is the degree of substitution, I Protein−methacrylate, I Protein and I Reference are the intensities of the fluorescent signal for the methacrylated protein, the unreacted protein and the reference, respectively. Hydrogel Fabrication Ferritin-GelMA, FerMA-GelMA, apoferritin-GelMA and ApoMA-GelMA were obtained by dissolving 5 wt% freeze dried GelMA and 0.5% (w/v) of the photoinitiator (PI), 2-hydroxy-1-(4-(hydroxyethoxy)-phenyl)- 2-methyl-1-propanone (Irgacure 2959) in PBS for 1 hour at 80 oC. Different concentrations of ferritin, FerMA, apoferritin and ApoMA were added to the solution and mixed. Then 30 μl pre-polymer solution was pipetted on an even surface with a spacer height of 600 μm. The drop was covered with a TMSPMA-coated glass slide, producing a cylindrical specimen with 600 μm height and 114 6mm diameter. Then it was exposed to UV light with the intensity of 6.9mW/cm2 and the distance of 8 cm for crosslinking. Evaluation of Hydrogel Mechanical properties: The hydrogel disks were compressed at 1mm min-1 until they fractured using a mechanical testing system, Instron 5942 Single Column Tabletop Model, an ITW company (Illinois Tool Works Incorporation, Glenview, Illinois, USA), equipped with a computer-based control/ analysis system (Bluehill Version 3). The compressive modulus was evaluated as the slope of the linear region corresponding with 0–15% strain of stress-strain curve. Ultimate stress was determined as the maximum stress before the hydrogel fractured. SEM Imaging SEM was used to analyze the morphological features of hydrogels. Hydrogels were washed with DI water and lyophilized. Then, the dried hydrogel sample were conducted with a Philips XL30S FEG with a SiLi (Lithium drifted) with a Super Ultra-Thin Window EDS detector and with a LEICA LEO S430i SEM. Cell Studies NIH-3T3 fibroblasts were suspended in a prepolymer solution (2*10 6 cell/ml) and crosslinked via photocrosslinking to fabricate hydrogels. The hydrogel laden cells were incubated in media (Dulbecco’s modified Eagle medium (DMEM) supplementing with 10% fetal bovine serum (FBS) and 100 units/mL penicillin-streptomycin, all purchased from Invitrogen) at 37 oC and 5% CO2. Viability was measured with a LIVE/DEAD® Viability/Cytotoxicity Kit. Samples were incubated with calcein (green) and ethidium homodimer-1 (red) at a dilution of 0.5:1000 and 2:1000 respectively for 10 minutes at 37 115 °C. Samples were imaged using an inverted fluorescent microscope (Zeiss, Axio Observer.D1, Germany), and ImageJ was used for the analysis of the images. On the 7th day of culture, samples were fixed with 4% paraformaldehyde at room temperature for 30 minutes. Following fixation, samples were permeabilized with 0.1% Triton X-100 for 30 minutes. The samples were stained by F-actin with Alexa Fluor 488 phalloidin at a 1:40 dilution for 45 minutes. Finally, the samples were counterstained with DAPI, which was prepared at a dilution of 1:1000 for 15 minutes and were imaged an inverted laser scanning confocal microscope (Leica SP5X MP, Germany) Fabrication of FITC laden Apoferritin hydrogel and characterization 1 mg/ml and 0.1 mg/ml apoferritin and ApoMA were prepared in the HCL solution with pH 2. The solution was kept overnight at 4 oC for disassembling the ApoMA and apoferritin. 20 µl of 1Mole NaOH was added to the solution afterward to neutral the solution. Afterwards, solution was dialyzed with dialysis cassette (Slide- A-Lyzer, MWCO 12000, Fisher Scientific) for 2 days at 4 oC. Finally, 5 wt% GelMA, 0.5 photoinitiator wad mixed with apoferritin loaded with 2 mg/ ml FITC solution (FITC solution was prepared in PBS) and crosslinked as explain in the fabrication of hydrogel part. Each FITC laden apoferritin hydrogel specimen (600 μm height and 6mm diameter) was placed in 500 µl PBS and accumulative release test was performed at different time points (0 hr, 6 hr, 12 hr, 24 hr, 48 hr, and 72 hr). Statistical analysis: Statistical data was analyzed using a one-way ANOVA (GraphPad Prism 5.02, GraphPad Software) when appropriate. The data represents the mean, and the error bars 116 represent ± standard deviation (SD) of each sample. Data was classified as significant if p < 0.05 following a Tukey’s multiple comparison test. B.3. Results and Discussion B.3.1. Synthesize and characterization of methacrylated ferritin and apoferritin Methacrylate groups were conjugated into ferritin and apoferritin by reaction with methacrylate anhydride (Figure B.1a and B.1b) to achieve the chemical bonding of ferritin and apoferritin with GelMA. To determine the conversion of amines, a fluoraldehyde assay was used. The fluorescent signals give direct information of free amine groups and therefore be an indicator of the percentage of the reacted amine groups. The results showed that the degree of methacrylation of FerMA is ~85% and ApoMA is ~92 % (Figure B.1c), confirming reasonable amine groups of ferritin and apoferritin conjugated with methacrylate groups. Furthermore, the structure integrity of the ferritin and apoferritin after methacrylation were verified using a transmission electron microscope (TEM) (Figure B. 1d). Comparing the TEM images of ferritin with FerMA and apoferritin with ApoMA, it could be concluded that both apoferritin and ferritin did not collapse during methachrylation process and retained their nanocage structure which is necessary for drug delivery potential. 117 Figure B.1. Characterization of methacrylated ferrtin/apoferritin. (a) Schematic of Ferritin methacrylation (b) Schematic of Ferritin methacrylation (c) Methacrylation degree of ferritin and apoferritin. (d) Transmission Electron Microscope (TEM) of (i) Ferritin, (ii) FerMA, (iii) Apoferritin, and (iv) ApoMA. Statistical analyzes have been performed, * p < 0.05. B.3.2. Morphological evaluation of nanoparticle composite GelMA hydrogels GelMA hydrogels incorporated with varying amounts of ferritin (‘ferritin-GelMA hydrogel’), FerMA (‘FerMA-GelMA hydrogel’), apoferritin (‘apoferritin-GelMA hydrogel’), ApoMA (‘ApoMA-GelMA hydrogel’) were fabricated by photo-crosslinking method (Figure B.1a and B.2a). The visual change in the color of ferritin-GelMA and apoferritin-GelMA pre-polymer solution when the concentration of ferritin/apoferritin increased was shown in Figure B.2b. Here, as the ferritin concentration was increased, the 118 pre-polymer solution showed a stronger brownish yellow color because of ferritin iron core. On the other hand, change in concentration of apoferritin, which do not have iron core, in the pre-polymer solution did not affect the visible properties and the solution remained colorless and transparent (Figure B.2b). To initiate gelation of pre-polymer solution through crosslinking of methacrylate groups, Irgacure 2959 photo initiator (PI) which was activated upon UV exposure in the range of 200nm – 300nm was used[199]. Considering ferritin/apoferritin has significant absorbance in that range, ferritin-GelMA and apoferritin-GelMA pre-polymer solution require longer polymerization time and UV exposure to form a gel compared to the prepolymer solution without ferritin and apoferritin[200]. The absorbance of the composite GelMA pre-polymer solution (without PI) was investigated for different concentration of ferritin, apoferritin, FerMA, and ApoMA (Figure B.2c and Figure B.6). By increasing the concentration, the absorbance was increased considerably, suggesting more UV exposure time is required for crosslinking of higher concentration of ferritin-GelMA, FerMA-GelMA, apoferritin-GelMA and ApoMA-GelMA hydrogels prepolymer. Due to its iron core, ferritin has a broad absorption band in the ultraviolet region which tails into the visible region of the spectrum. However, a significant part of the measured absorbance of ferritin would be attributed to the scattering from the protein shell rather than absorbance by the core[200]. Hence, despite the lack of iron core, the measured UV absorbance of apoferritin was still very high. To obtain more insight into the effect of covalently incorporating ferritin into GelMA hydrogel, scanning electron microscopy (SEM) was used to compare the porosity and morphology of pure GelMA and ferritin/FerMA incorporated GelMA hydrogels. By adding 0.1 mg/ml ferritin or FerMA, the composite hydrogels showed increased pore size, 119 as well as rougher and structured surface, suggesting a beneficial effect of integrated nanoparticles on microstructure of GelMA hydrogel (Figure B.2d). The complex and structured morphology of hydrogel may contribute to strengthen structural and mechanical robustness, as well as improved surface topographical cues of cells, which are crucial for tissue engineering applications [201, 202]. SEM images of GelMA-FerMA showed a more structured pore compare to GelMA-ferritin. This could be due to bonding between FerMA and GelMA, where FerMA acted as a spring in GelMA network and therefore, created more robust structures. In addition, the smooth surface of the pore walls of incorporated hydrogel confirmed that both ferritin and FerMA were homogeneously dispersed in GelMA without severe aggregation. Figure B.2. Synthesis and characterization of ferritin/apoferritin-GelMA hydrogels. (a) Schematic of Ferritin methacrylation with GelMA (b) Optical images of ferritin-GelMA 120 and apoferritin-GelMA prepolymer with different concentration of ferritin and apoferritin (c) UV-Vvis absorption spectra of ferritin-GelMA and apoferritin-GelMA prepolymer solutions with different concentrations of ferritin and apoferritin, grey range (320 nm-390 nm) indicate the wavelength of UV curing system. (d) Scanning electron microscopy (SEM) of 5 wt% GelMA only, 5 wt% GelMA- 0.1 mg/ml ferritin, and 5 wt% GelMA- 0.1 mg/ml FerMA. Scale bar is 150 nm. B.3.3. Mechanical Properties of nanoparticle composite GelMA hydrogels One of the significant challenges in polymer hydrogel fabrication for tissue engineering applications is the realization of materials with a high stiffness. Specially, fabrication of GelMA which is a porous hydrogel structures with fast water absorption require high mechanical properties, i.e., the elastic modulus, tensile strength, and toughness. To evaluate the effects of covalently combined ferritin on the mechanical properties of GelMA scaffolds, mechanical properties were tested on the various concentrations of ferritin, apoferritin, FerMA, and ApoMA incorporated GelMA. Initially, the effect of UV crosslinking time on Young’s modulus of GelMA with ferritin, apoferritin, FerMA, and ApoMA composites were investigated (Figure B.3a, B.3d). Considering UV absorbance of Ferritin, apoferritin, FerMA, and ApoMA (discussed before), the UV exposure time is an important factor affecting mechanical properties of hybrid hydrogels. The optimization of curing time is essential not only for tuning mechanical properties but also minimizing damage that the UV irradiation can cause to encapsulated cells. Young’s modulus of 5 wt% GelMA with different concentration of ferritin were measured for various UV exposure time (Figure B.3a). As a general trend in GelMA with different 121 ferritin concentrations, the Young’s modulus was positively correlated to the UV exposure time and it reached a plateau followed by a slight downward trend after about 60 sec of exposure. This is to be expected as the hydrogels become fully crosslinked when sufficiently exposure to UV irradiation and therefore further exposure could not enhance the mechanical properties. Young’s modulus for 5 wt% GelMA starts at 2.36 kPa for 10 sec UV exposure and reaches it maximum at 6.24 kPa for 60 sec UV exposure. Adding 0.1 mg/ml ferritin to GelMA significantly enhanced the Young’s modulus of hydrogel where the hydrogels have ~3 times higher Young’s modulus than that of GelMA only. Nanocomposites composed of functional nanoparticles that are embedded within polymer hydrogels have begun to attract much attention as a new generation of biomaterials because of their tunable mechanical properties. In addition to the intrinsic mechanical properties of both the hydrogel and incorporated nanoparticle, there are still three important parameters deciding the mechanical behavior of the nanocomposites, including the nanoparticle architecture, nanoparticle-hydrogel and nanoparticle–nanoparticle interactions[203]. Regarding to ferritin, particle-hydrogel interactions play a crucial role in the degree of mechanical reinforcement in nanocomposites. It was proposed that the functional groups on protein shell of ferritin (carboxylic acid (–COOH) and amino (–NH2) groups) could combined with GelMA (hydroxyl (–OH) group) by forming hydrogen bonds[195]. Therefore, the protein shell act as an elastic nanosprings between the GelMA and ferritin core to improve the mechanical properties of ferritin/FerMA-GelMA. However, when further increase of ferritin concentration to 1 mg/ml, the Young’s modulus of hydrogel drops significantly and becomes similar to that of pristine GelMA. Moreover, for higher ferritin concentration (10 mg/ml) in GelMA, gelation was not achieved when exposure 122 time is less than 180 sec. It is concluded that excess ferritin acting as barrier between GelMA fibers and preventing them from creating an interconnect network because of the limited hydroxyl groups on GelMA. Meanwhile, due to the increased iron particle density, higher concentration of ferritin absorbs a high volume of UV light, resulting the decrease of Young’s modulus[200]. In contrast, 1 mg/ml FerMA-GelMA further improved the Young’s modulus of the composite hydrogel to ~4 time higher than that of GelMA only when compared with 0.1 mg/ml FerMA-GelMA. This could be attributed to the different functional groups involved in the bond between FerMA and GelMA compared with ferritin and GelMA. Exposure of UV light could initiate the chain polymerization of the methacryloyl substitutions between FerMA and GelMA. Incorporated FerMA do not compete for hydroxyl (–OH) group on GelMA, and therefore, more FerMA could further improve mechanical stiffness without hindering the interconnected network formation during GelMA crosslinking despite the high UV absorption. Similar study was performed on 5 wt% GelMA hydrogels with different concentrations of apoferritin and ApoMA (Figure B.3d). Here, for 0.1 mg/ml apoferritin hydrogel, the Young’s modulus was enhanced by ~2 folds compared to GelMA only. Young’s modulus increase of apoferritin-GelMA was still lower compared to improvement produced in the hydrogel samples by adding 0.1 mg/ml ferritin to GelMA (Figure B.3a), which could be the result of a missing iron core in apoferritin[194]. By further increase of apoferritin concentration to 1 mg/ml, the Young modulus dropped significantly due to the stronger scattering from the protein. Also, in the case of adding 0.1 mg/ml and 1mg/ml ApoMA, the hydrogel mechanical properties were improved by 2 folds compared to 123 GelMA only with no significant difference between two concentrations because of same reasons of 1 m/ml FerMA-GelMA hydrogel. Young’s modulus of GelMA with different concentration of ferritin/FerMA is compared with each other for 10 sec UV exposure (Figure B.3b and B.3c). Minimum exposure time can be important to minimize light-toxicity when the hydrogels contain cells [204]. Here, 10 sec is chosen as the minimum exposure time when gelation was observed in all hydrogels. The evaluation was performed for GelMA incorporated with 0 mg/ml, 0.1 mg/ml, and 1 mg/ml of ferritin/FerMA (Figure B.3c). 0.1 mg/ml ferritin/FerMA-GelMA has the maximum young modulus at 10 sec UV exposure (Figure B.3c). By increasing the concentration of FerMA to 1 mg/ml, the Young’s modulus decreased as expected, which may be attributed to higher portion of UV absorption (Figure B.2c). Due to the short time of UV exposure, higher concentration of FerMA could not offset the effect of UV absorption which attenuated the chain polymerization of the methacryloyl substitutions between FerMA and GelMA, and between GelMA polymers. As a result, the hydrogel would have less degree of crosslinking and corresponding lower Young’s modulus in 1 mg/ml FerMA-GelMA. In the case of GelMA with apoferritin and ApoMA, Young’s modulus showed similar trend. The tensile tests for different concentrations of ferritin/FerMA-GelMA are shown in Figure B.3g. Under tensile test, ferritin/FerMA composite hydrogels showed linear relation between stress and applied strain. In terms of ultimate tensile strain (maximum elongation), approximately 20% improvement was observed for ferritin/FerMA-GelMA with one exception. In the case of GelMA with 1 mg/ml ferritin, ultimate tensile strain slightly decreased compared to GelMA alone. Moreover, ultimate tensile stress increased 124 for all ferritin/FerMA cases compared to GelMA only. Here, about 100% and 60% improvement in ultimate tensile stress was observed for 0.1 mg/ml ferritin/FerMA-GelMA and 1 mg/ml ferritin/FerMA-GelMA, respectively (Figure B.3g). Similar trend was observed when different concentrations of apoferritin/ApoMA were added to GelMA hydrogels (see Figure B.3h). Smilarly, addition of apoferritin/ApoMA to GelMA hydrogels increased the ultimate tensile strain by approximately 20% for all composites with only one exception, where addition of 1 mg/ml apoferritin to GelMA reduced the ultimate tensile strain. In this case, the composite hydrogel shows the lowest ultimate tensile stain compared to other cases. Also, addition of apoferritin/ApoMA to GelMA hydrogels improved the ultimate tensile stress of the composite hydrogels, where 0.1 mg/ml and 1 mg/ml apoferritin, and 1 mg/ml ApoMA show approximately 100% improvement and 0.1 mg/ml ApoMA shows about 50% improvement compared to GelMA only. Overall, addition of methacrylate form of ferritin/apoferritin (FerMA/ApoMA) to GelMA hydrogel significantly improved both maximum elongation (ultimate tensile strain, Figure B.3k and l) and ultimate tensile stress of the composite hydrogel (Figure B.3j). However, in the case ferritin/apoferritin, improvement in both ultimate strain and ultimate stress was observed merely for 0.1 mg/ml concentration, and further increase in the concentration of apoferritin/ferritin to 1 mg/ml decreased maximum elongation of the hydrogel to the level of GelMA only (Figure B.3j). To further specify the hydrogel, the toughness of different concentrations of nanoparticle incorporated GelMA are evaluated. As shown in Fig. 3i, toughness for 5 wt% GelMA was 343.2 J/mm3. Here, adding ferritin/FerMA/apoferritin/ApoMA significantly enhanced toughness of hydrogel in all cases. However, the toughest composite hydrogels 125 were achieved by incorporating 0.1 mg/ml ferritin/FerMA/apoferritin/ApoMA, where the toughness was enhanced by more than 135%. The corresponding values are 885.1 J/mm3, 809.8 J/mm3, 864.3 J/mm3 and 831.2 J/mm3 for 0.1 mg/ml ferritin, 0.1 mg/ml FerMA, 0.1 mg/ml apoferritin and 0.1 mg/ml ApoMA, respectively. For higher concentration of nanoparticles (1 mg/ml), the toughness slightly decreased in all cases compared with 0.1 mg/ml. However, the lowest value of toughness obtained was 600.5 J/mm3 for 1 mg/ml ferritin, which was still 75% higher than GelMA only. As described above, Shin et. al. found that the elastic modulus, mechanical strength and maximum elongation of the PVA/ferritin hydrogel increased significantly compared to pristine PVA. It was proposed that since the protein shell of ferritin is comprised of elastic-like, peptidic alpha-helical bundles, the ferritin/apoferritin acts as a spring and allow consistent nanoparticle-hydrogel interaction during hydrogel compression and elongation, which could also explain the reinforcement of tensile stress and strain in ferritin/GelMA hydrogels. Detailly, the mechanism underly the enhancement of mechanical properties in the ferritin/GelMA can be concluded that the protein shell with a stiffness higher than that of the GelMA matrix can reduce tensile hoop stress at the hole boundary of incorporated reinforcing material[194]. Because of the relationship of E iron core>E protein shell>E GelMA (E denotes elastic modulus), the elastic modulus of the ferritin/apoferritin-GelMA hydrogels increased significantly. Overall, the interaction of the ferritin with the GelMA matrix, as well as the relationship of elastic modulus, influence the enhancement in toughness of the nanocomposite GelMA hydrogels. Therefore, in our study, the nanoparticle concentration and UV exposure time which could affect the distance between arranged particles and the 126 bonding connection between particle and GelMA matrix are the most important factors of determining the mechanical reinforcement of nanocomposite hydrogels. 127 Figure B.3. Strengthen mechanical properties of nanoparticle incorporated composite hydrogels. (a) Young modulus of GelMA with different concentrations of ferritin and FerMA versus UV exposure time (b) Stress-strain curves for compression test of GelMA with different concentration of ferritin and FerMA (UV exposure time 10 sec) (c) Young 128 modulus GelMA with different concentrations of ferritin and FerMA (UV exposure time 10 sec) (d) Young modulus of GelMA with different concentrations of apoferritin and ApoMA versus UV exposure time (e) Shear-stress of GelMA with different concentration of apoferritin and ApoMA (UV exposure time 10 sec) (f) Young modulus GelMA with different concentrations of apoFerritin and ApoMA (UV exposure time 10 sec) (g) Stress-strain curves for tensile tests of GelMA with different concentration ferritin and FerMA (h) Stress-strain curves for tensile tests for GelMA with different concentration apoferritin and ApoMA. (i) Toughness of GelMA with different concentration ferritin, FerMA, Apoferritin and ApoMA. (j) Calculated tensile stress vs maximum elongation. (k) Images of hydrogel under tensile test. Statistical analyzes have been performed, (data set that show not significantly difference (ns. ;> 0.05), data set that are at significant levels: * p < 0.05, ** p < 0.01, *** p < 0.001. B.3.4. Cell encapsulation in nanoparticle composite GelMA hydroels It was reported that the mechanical properties of the scaffold matrix including the porosity are recognized to affect cellular behavior such as proliferation, elongation and differentiation [170, 181, 201]. Hence, it is of great important to maintain their high biocompatibility advantage while tuning mechanical robustness for tissue engineering application. Here, the pre-polymer solution was exposed to 10 sec UV for gelation based on the results of mechanical properties. 0.1 mg/ml and 1 mg/ml ferritin, FerMA, apoferritin, ApoMA with 5 wt% GelMA were investigated in the study and the cell viability and elongation potential of encapsulated NIH-3T3 fibroblasts was evaluated. As shown in Figure B.4a and Figure B.6, the difference in cell viability after 1 day of culture between 129 pristine GelMA and GelMA with nanoparticles was not significant. It was elucidated experimentally that the cell viability was higher than 95% for all samples (Figure B.4b), proving biocompatibility of the incorporated nanoparticles, even after methacrylation. Phase contrast images showed that encapsulated fibroblasts in 5 wt% GelMA incorporated with 0.1 mg/ml, 1 mg/ml ferritin and FerMA could be observed cell spreading as soon as on day 2 whereas in the case of GelMA only (Figure B.7), cell spreading did not show on day 2. Percentage of spreading fibroblasts in 5 wt% GelMA with different concentration of ferritin and FerMA are quantified in Figure B.4d for two different UV exposure time; 10 sec and 20 sec. The results showed that cell spreading percentage in GelMA with ferritin/FerMA are higher than GelMA only. Interestingly, cell spreading percentage of GelMA with 0.1 mg/ml ferritin for 10 sec and 20 sec UV exposure was almost the same. However, for GelMA with 1 mg/ml ferritin, 0.1 mg/ml FerMA, and 1 mg/ml FerMA, the percentage of spreading cell significantly decreased for 20 sec UV exposure. Therefore, 10 sec UV exposure time was found to be optimal for encapsulated cell spreading. 130 Figure B.4. Viability and elongation of cells encapsulated in nanoparticle incorporated composite hydrogels. (a) Live dead assay images (b) Quantification live/dead (c) Phase contrast images 3T3 cell encapsulated in GelMA -ferritin and GelMA- FerMA, for 10 sec 131 UV exposure time at day 2 of culture. (d) Quantification chart of speaded cells in GelMA -ferritin and GelMA- FerMA for 10 sec and 20 sec UV exposure time at day 2. (e) Actin/DAPI staining of 3T3 fibroblast encapsulated in 5 wt% GelMA hydrogels with different concentration of ferritin and FerMAafter 7 days of culturing (UV exposure time 10s). DAPI (blue) for nuclei, phalladoin (green) for f-actin. Images taken with 100x magnification, white scale bar equals to 200 μm, (f) Percentages of actin covered with filamentous actin. (g) Actin/DAPI staining of 3T3 fibroblast encapsulated in 5 wt% GelMA hydrogels with different concentration of apoferritin and ApoMA after 7 days of culturing (UV exposure time 10s). DAPI (blue) for nuclei, phalladoin (green) for f-actin. Images taken with 100x magnification, white scale bar equals to 200 μm, (h) Percentages of actin covered with filamentous actin. Statistical analyzes have been performed, (data set that show not significantly difference (ns. ;> 0.05), data set that are at significant levels: * p < 0.05, ** p < 0.01, *** p < 0.001. As a key indicator for cell spreading and cell motility, the filamentous actin of the cells inside hybrid hydrogels after 7 days of culture was evaluated to display elongated, proliferated, and well-interconnected cellular shapes (Figure B.3e and g). Here, for GelMA with ferritin, FerMA, and apoferritin, more actin filament networks were observed when compared with GelMA only. It was also shown that the cells were more elongated in the case of 0.1 mg/ml more than 1 mg/ml FerMA/ApoMA-GelMA hydrogels (Figure B.3f and h). Consistent with our findings, several previous studies have also reported the enhanced cell behavior, e.g. higher cellular attachment and retention, when nanoparticles are incorporated in the hydrogel[181, 186, 205]. Such behavior could be explained that 132 increased surface roughness in nanoscale, which are attributed to the presence of nanomaterial (ferritin, FerMA, apoferritin and ApoMA) within GelMA hydrogel, can improve surface topographical cues for culturing through integrin receptors[202]. Roughness, which relates to the texture of the uppermost layer of a material, is the most investigated aspect of topography. Abundant experiments have reported that surface roughness can influence cell behaviors like adhesion, migration, proliferation and differentiation. Therefore, ferritin/apoferritin promoted cell viability and elongation without inducing cytotoxicity during the 7 days of culture, validating the feasibility of utilizing composite hydrogels with beneficial cellular microenvironment which plays a crucial role in improving cell response and function for engineering tissue application. B.3.5. FITC release from GeMA-Apoferritin and GelMA-ApoMA encapsulated FITC Recently, nanoparticle is one of the most investigated ways for effective drug delivery. These nanocarriers are characterized by sustained release, targeted delivery, and dose control. Among many different nanoparticles which have been developed for drug delivery including polymeric nanoparticles, liposomes and nanochitosan, apoferritin has merged as an excellent protein-based nanocage thanks to its unique structure, surface properties and high biocompatibility. With these regards, in the present study, FITC laden apoferritin/ApoMA-GelMA were designed and assessed to show the drug delivery characteristics of these composite hydrogels. Apoferritin can reversibly dissociate and associate in a shape memory fashion based on the surrounding pH. It has been proposed that while disassembled, apoferritin can be 133 mixed with drug molecules, and drug molecules are encapsulated within the apoferritin cavity once apoferritin is reassembled[190, 206]. In this study, we investigated drug delivery potential of apoferritin-GelMA by detecting the release profile of encapsulated FITC. As shown in Fig 5a, FITC was encapsulated in apoferritin as describe in previous studies for encapsulation of doxorubicin in apoferritin[207]. To validate the nanocage function of apoferritin/ApoMA, absorption spectrum of FITC only solution, filtered FITC mixed with apoferritin (FITC was not encapsulated), filtered FITC encapsulated apoferritin and apoferritin were measured. The apoferritin showed high absorbance in UV spectrum with a specific absorbance peak of aromatic amino acids at 280 nm (Figure B.5b). Apoferritin mixed with FITC only showed absorbance peaks for apoferritin. Differently, FITC encapsulated apoferritin showed absorbance spectrum containing both peaks characteristic for FITC and apoferritin, implying the successful encapsulation of FITC in protein cage. Our results confirmed that there are few unspecific bonds between FITC and apoferritin surface which was further validated by the fluorescent images (Figure B.5c), providing evidence to the reliability of FITC release method. To evaluate the cumulative release profile, the release test was performed on apoferritin/ApoMA-GelMA hydrogels with encapsulated FITC at different time points (0 h, 6 h, 12 h, 24 h, 48 h, and 72 h). The results revealed that ApoMA-GelMA and apoferritin-GelMA have similar sustain-release profiles (Figure B.5d). Here, during the first 24 h, FITC was released at a constant rate and the release profile slope is constant. The release rate, however, gradually slowed down after 24 h. Eventually, almost 90% of the FITC had been released after 72 h in both ApoMA-GelMA and apoferritin-GelMA cases. Features 134 of apoferritin have been extensively studied by researchers, contributing to the use of apoferritin as biological nanoparticles for nanomedical applications. To take advantages of apoferritin nanocages, apoferritin/ApoMA conjugated GelMA hydrogels were designed and fabricated in our study for sustained drug delivery and tissue engineering application. Our results showed that apoferritin/ApoMA-GelMA maintain the structural integrity and intact function of nanocages. Taken together, apoferritin/ApoMA-GelMA will be a promising candidate of drug laden hybrid hydrogels for biomedical applications. A longer period of sustained and controllable drug release which could be attained by optimizing the hydrogel and nanoparticle characteristics is crucial for clinical transplants. Hence, further studies are needed to optimize release profile of apoferritin/ApoMA-GelMA hydrogels. 135 Figure B.5. Drug delivery potential of apoferritin/ApoMA-GelMA hydrogels. (a) Schematic of capturing FITC with Apoferritin. (b) Absorbance spectra of plain FITC (red color), Filtered FITC captured in Apoferritin (blue color), Filtered FITC mixed with Apoferritin (green color), and Apoferritin (black color) were measured. (c) Florescence images of (i) Plain FITC, (ii) Filtered FITC captured in Apoferritin, and (iii) Filtered FITC mixed with Apoferritin. Scale bar is 2mm. (d) FITC release profile of FITC laden GelMA-Apoferitin and GelMA-ApoMA. 136 B.4. Conclusion In this study, ferritin/apoferritin was successfully methacrylated in order to covalently bond with a GelMA hydrogel network by photo-crosslinking method. The structure of ferritin/apoferritin remained preserved during the methacrylation process as confirmed by TEM images. Methacrylation of ferritin and apoferritin showed better incorporation (of higher concentration) in the GelMA network. Incorporation of ferritin, FerMA, apoferritin, and ApoMA with GelMA hydrogels successfully improved the mechanical properties including elastic modulus, maximum elongation and toughness. The addition of ferritin created larger interconnected pores and rougher surface of hydrogel, resulting in better cell spreading and elongation. Furthermore, the biocompatibility of ApoMA/FerMA hydrogels were confirmed by evaluating the viability of 3-D encapsulated NIH-3T3 fibroblast cells. The cell studies indicated that cells spread faster and elongated more in the nanocomposite hydrogels compared to pristine GelMA hydrogel. Moreover, the application of apoferritin/ApoMA for drug delivery was confirmed by FITC release test where FITC cumulative release profile was assessed over 70 hours from FITC loaded apoferritin-GelMA hydrogel. The results thus confirm that the developed apoferritin/ApoMA-GelMA hydrogels can be used as a nanocarrier for sustained drug release and a strong stiffness hydrogel for tissue engineering applications. In one possible application, these tuned hydrogels can also be used as an injectable biomaterial for sustained drug delivery in post-surgical tumor treatment. 137 Figure B.6. UV-Vvis absorption spectra of FerMA-GelMA and ApoMA-GelMA prepolymer solutions with different concentrations of FerMA and ApoMA, grey range (320 nm-390 nm) indicate the wavelength of UV curing system. Figure B.7. Actin/DAPI staining of 3T3 fibroblast encapsulated in 5 wt% GelMA hydrogels with 10 sec UV exposure time after 7 days of culture. DAPI (blue) for nuclei, phalladoin (green) for f-actin. Images were taken with 100x magnification, 138 Figure B.8. Phase contrast images of 3T3 fibroblast encapsulated in ferritin-GelMA and FerMA-GelMA with 10 sec UV exposure time after 2 days of culture. 139 Appendix C G-Code for cellink bioprinter G90 ; ABSOLUTE G21 ;metric values G91 ;RELATIVE positioning ;start the code ;-------------------------------- ;Printing block 00 ;-------------------------------- ;Printing 1st row G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 140 G1 Z7 F120 ;--------------------------------- ;Printing 2nd row ; y movement G1 Y-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------ ;Printing 3rd row 141 ; y movement G1 Y-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------------------ ;-------------------------------- ;Printing block 01 ; y movement G1 Z-2 F600 142 G1 Y-3.25 F120 ;--------------------------------- G4 P10000 ;Printing 1st row G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------ ;Printing 2nd row ; y movement G1 Y-0.45 F600 G1 Z-2 F600 143 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------------------- ;Printing 3rd row ; y movement G1 Z-2 F600 G1 Y-0.45 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 144 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------------------------ ;------------------------------------ ;Printing block 11 ;-------------------------------- ; x movement G1 Z-2 F600 G1 X4.5 F600 ;-------------------------------- G4 P10000 ;Printing 1st row G1 Z-5 F600 145 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;--------------------------------- ;Printing 2nd row ; y movement G1 Y0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 146 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------ ;Printing 3rd row ; y movement G1 Y0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 147 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------------------- ;-------------------------------- ;Printing block 10 ; y movement G1 Z-2 F600 G1 Y3.25 F600 ;--------------------------------- G4 P10000 ;Printing 1st row G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 148 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------ ;Printing 2nd row ; y movement G1 Y0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 149 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------------------- ;Printing 3rd row ; y movement G1 Y0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 150 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ; x movement G1 X-0.45 F600 G1 Z-2 F600 ;z inside the gel G1 Z-5 F600 G1 Z7 F120 ;------------------------------------ G1 Z7 F700 ;----------- G0 X0 Y0 Z2 F600 G91 G90 M84 151 Bibliography 1. 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Thesis/Dissertation
eng
Mechanical Engineering
Vancouver : University of British Columbia Library
Attribution-NonCommercial-NoDerivatives 4.0 International
Graduate
Development of 3D functional brain tissue model
Text