Le développement et la modélisation numérique d'un bioréacteur pour l'ingénierie des tissus de grande masse

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UNIVERSITY DE

SHERBROOKJE

Faculte de genie

Departement de genie chimique et genie biotechnologique

LE DEVELOPPEMENT ET LA MODELISATION NUMERIQUE

D'UN BIOREACTEUR POUR L'INGENIERIE DES TISSUS DE GRANDE MASSE

DEVELOPMENT AND NUMERICAL MODELING OF A BIOREACTOR

SYSTEM FOR THE ENGINEERING OF LARGE-SCALE TISSUE

These presentee pour l'obtention du grade de Philosophise Doctor (Ph.D) en sciences appliquees

Speciality : genie chimique

Davod Mohebbi Kalhori

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Composition du jury /Doctoral Committee:

Prof. Charles Doillon

Prof. Amin Behzadmehr

Prof. Bernard Marcos

Prof. Yves Mercadier

Prof. Gerard Lachiver

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In memory of my loved and dearly-missed father and teacher

This work is dedicated to those I love dearly:

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RESUME

L'ingenierie tissulaire comme domaine multidisciplinaire combine les principes et les metho-des de l'ingenierie et metho-des sciences de la vie pour resoudre les problemes biomedicaux dans metho-des applications cliniques. Toutefois, le defi majeur consiste a trouver une maniere de faire passer le produit d'une echelle de recherche a une plus grande echelle de production, en permettant que les tissus fonctionnels construits soient reproductibles, securitaires et economiquement competitifs. Pour resoudre ces problemes, les bioreacteurs ont ete mentionnes comme essen-tiels afin d'ameliorer l'ingenierie des tissus vivants in vitro. Pour ces raisons, l'objectif general de cette these est le developpement d'un bioreacteur avance et des methodes numeriques et de visualisation etant deux precieux outils pour mieux concevoir ce bioreacteur et ainsi compren-dre les mecanismes. Ces mecanismes incluent les processus ainsi physiques, chimiques et bio-logiques dans un environnement trois dimensions (3-D), conduisant a faciliter la production, a grande echelle, de tissus fonctionnels in vitro.

Pour atteindre ces objectifs, la presente these comprend une etude experimentale et une etude de modelisation numerique, qui se deroulent en quatre etapes distinctes : 1) La conception et la construction d'un bioreacteur, ainsi que 1'evaluation et le controle de son hydrodynamique. 2) la visualisation de la perfusion de l'ecoulement du fluide dans le bioreacteur a membrane a fibres creuses (HFMB) en utilisant la technique d'imagerie biomedicale noninvasive (i.e. la tomographie par emission de positrons (TEP)). 3) le developpement d'un modele mathemati-que pour l'analyse du bioreacteur a membrane a fibres creuses hybride (hHFMB) et 4) le de-veloppement d'un modele dynamique et TWO-POROUS-MEDIA pour analyser le HFMB a l'aide de la dynamique des fluides computationnels (CFD), specifiquement pour les applica-tions en ingenierie des tissus osseux.

La partie experimentale comprend l'etape 1 et 2. Dans l'etape 1, le bioreacteur a perfusion a ete con?u et construit. L'evaluation de 1'hydrodynamique et du controle a ete effectuee. Dans ce systeme, la pression moyenne, le debit moyen, la frequence ainsi que la forme d'onde de la pression et de l'ecoulement pulsatiles peuvent etre modules et controles avec le temps pour simuler des conditions physiologiques et non physiologiques. La temperature, l'oxygene dis-sous et le pH peuvent etre controles. Ce systeme de bioreacteur peut etre applique a une varie-te de configurations, de geometries et de tailles de l'echafaudage du fait que le bioreacvarie-teur

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lui-meme est reglable en longueur. Ce systeme est autoclavable, et compatible avec des techni-ques d'imagerie noninvasive medicale. La conception des ports d'entree et de sortie du bio-reacteur ont ete realise par simulation CFD en regardant la distribution de l'ecoulement, pour atteindre un niveau d'efficacite elevee dans l'uniformite de la perfusion de l'ecoulement.

Dans la deuxieme etape, la TEP a ete proposee pour la premiere fois pour obtenir des nouvel-les informations sur nouvel-les modes d'ecoulement permanent et pulse dans le HFMB pour nouvel-les appli-cations en ingenierie tissulaire. Un petit systeme TEP pour animaux a ete utilise. La reparti-tion non-homogene du traceur, telle que trouvee avec la technique TEP, implique l'apparireparti-tion de regions non-efficaces en ce qui concerne le transfer! de masse. Pour la condition d'ecoulement permanent a l'entree, une non-homogeneite de la distribution des traceurs ra-dioactifs a ete obtenue. En revanche, l'ecoulement pulse a l'entree genere de la perfusion uni-forme mieux que pour un ecoulement permanent. En outre, il a ete trouve que pour les mernes conditions, l'accumulation du traceur dans le bioreacteur est plus faible pour l'ecoulement pul-satile que pour le permanent. Par consequent, ces resultats montrent que le TEP peut ameliorer la conception du bioreacteur, et ainsi ouvrir de nouveaux axes de recherche en ingenierie tis-sulaire.

La partie numerique comprend les etapes 3 et 4 dans lesquelles une etude numerique a ete re-alised pour le tissu osseux le HFMB comme etude de cas pour la culture tissulaire a grande echelle. Dans l'etape 3, la possibility d'utiliser le nouveau hHFMB pour la croissance des cel-lules souches mesenchymateuses (CSM) pour former le tissu osseux a ete etudiee en utilisant des simulations numeriques. Pour atteindre ce but, un modele mathematique utilisant un code CFD a ete concu pour optimiser les parametres de conception et de fonctionnement du hHFMB pour la croissance des CSM. La methode de volume moyen a ete utilisee pour formu-ler le bilan massique pour les elements nutritifs et les cellules dans l'espace extra-capillaire poreux du hHFMB. La cellule-echafaudage dans l'espace extra-capillaire des fibres creuses et le mur de membrane ont ete traites comme milieu poreux. La porosite, la permeabilite et la diffusivite qui dependent de la fraction volumique des cellules, ont ete utilisees dans ce mo-dele. Ces simulations ont permises la prediction simultanee de la distribution des nutriments et de la fraction volumique des cellules qui depend des elements nutritifs. De plus, ce modele a ete utilise pour etudier les effets des parametres de conception et de fonctionnement sur la dis-tribution des elements nutritifs et la croissance cellulaire. Les resultats de la modelisation ont

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demontre que la dynamique des fluides au sein de l'espace extra-capillaire et les proprietes de transport et les taux de consommation dans le hHFMB etaient suffisantes pour soutenir des CSM necessaires pour la production du tissu osseux en echelle clinique in vitro et permettre de resoudre les difficultes de nutrition en raison de la forte densite de la cellule de la taille de l'echafaudage.

Dans l'etape 4, le nouveau modele TWO-POROUS-MEDIA, afin de determiner la croissance cellulaire qui depend des elements nutritifs, a ete utilise pour analyser la croissance des CSM pour la formation de tissu osseux dans le HFMB. Dans ce modele, l'echafaudage en fibres creuses dans le bioreacteur a ete traite comme un domaine poreux. Le domaine se compose de la region lumen poreuse ou le fluide s'ecoule et la region ECS poreuse, remplie de gel de col-lagene contenant des cellules, afin de permettre la croissance de masse des tissus. En outre, les contributions de plusieurs parametres de processus et de conception, qui ameliorent les per-formances de bioreacteur, ont ete etudiees. En plus, revaluation dynamique de la croissance cellulaire et les distributions de l'oxygene et du glucose ont ete quantitativement analysees. Ces informations peuvent etre utilisees afin d'ameliorer la conception et les conditions opera-tionnelles du bioreacteur.

Dans cette etude, l'idee d'utiliser le TEP pour la visualisation des perfusions des ecoulements permanents et pulsatiles dans le HFMB proposee pour la premiere fois est tout a fait originale. En outre, tel qu'il a ete demontre dans la revue de la litterature, la combinaison des elements suivants n'a pas encore ete utilisee dans un bioreacteur pour des applications d'ingenierie tis-sulaire : 1) le developpement du HFMB permet l'ajustement de ses dimensions et est compa-tible avec 1'imagerie par resonance magnetique (IRM) et le TEP, 2) le bioreacteur plat, 3) la combinaison du HFMB avec un systeme pneumatique utilisant des tubes effondres pour pro-duire des ondes de forme pulsatile et 4) la surveillance et le controle du systeme. De plus, le hHFMB propose pour des applications en ingenierie du tissu osseux a I'echelle clinique et la modelisation du hHFMB demontre une fois de plus l'originalite de cette these. Enfin, toujours selon la revue de la litterature effectuee, le developpement du modele TWO-POROUS-MEDIA pour analyser la croissance des CSM pour la formation de tissu osseux dans le HFMB a ete propose et realise pour la premiere fois.

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Keywords: Bioreacteur a membrane a fibres creuses; Ingenierie tissulaire; Tissus de grande

masse; Ecoulement pulse; Tomographic par emission de positrons; modelisation numerique; Tissu osseux; Cellules souches mesenchymateuses

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ABSTRACT

Tissue engineering, as a multidisciplinary field, combines the principles and methods of engi-neering and life science in order to solve biomedical problems in clinical applications. How-ever, one major concern is how to scale up the products from research-scale into large-scale production of reproducible, safe, and economically competitive functional tissue constructs. To solve these problems, bioreactors have been shown to be essential in order to improve the

in vitro engineering of living tissues. To this aim, the general objective of this thesis is

devel-opment of an advanced bioreactor systems, visualization and computational methods that can serve as two valuable tools to facilitate better designing such bioreactor and understanding of the underlying mechanisms governing physical, chemical and biological processes in a three-dimensional (3-D) tissue culture environment, leading to facilitate functional large-scale tissue production in vitro.

To achieve the above mentioned general goals, this present thesis comprise two major parts both experimental and numerical study which have been conducted in four distinct steps as following: 1) Design, construction, and evaluation of control and hydrodynamic of a bioreac-tor system. 2) Visualization of fluid flow perfusion in the hollow fibre membrane bioreacbioreac-tor (HFMB) using a biomedical noninvasive imaging technique, i.e. positron emission tomogra-phy (PET). 3) Development of a mathematical model for analyzing a hybrid hollow fibre membrane bioreactor (hHFMB) and 4) Development of a dynamic and two-porous media model for analyzing the HFMB with the aid of computational fluid dynamics (CFD), specifi-cally for bone tissue engineering application.

The experimental part includes the steps 1 and 2. In the step 1, the flow perfusion bioreactor system has been designed and constructed. The experimental evaluations of hydrodynamic, and control were performed. In this system, mean pressure, mean flow rate, frequency and waveform of the pulsatile pressure and flow rate can be modulated and controlled over the time to simulate both physiological and non-physiological conditions. The temperature, dis-solved oxygen, and pH can be controlled. This bioreactor system can be applied to a variety of scaffold configurations, geometries, and sizes as the cell/tissue culture chamber is adjustable in length. This system is autoclavable, and compatible with noninvasive medical imaging techniques. Designing of the inlet and outlet manifold of the bioreactor were performed

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according to data obtained from CFD simulation of the flow distribution to achieve high effi-ciencies in the uniformity of flow perfusion.

In the second step, PET was proposed for the very first time and a small animal PET system was used to obtain new information about steady and pulsatile flow patterns in the HFMB for tissue engineering applications. The non-homogeneous tracer distribution, as found with PET imaging, implies the occurrence of non-efficient regions with respect to mass transfer. In steady inlet flow condition, a non-uniform distribution of radioactive tracer was obtained. In contrast, the pulsatile inlet flow generated more uniform perfusion than that of steady flow. Further, it was found that in the case of pulsatile flow, the accumulation of the tracer within the bioreactor was efficiently less than that of steady inlet flow at the same condition. There-fore, in one hand these findings have the potential to improve bioreactor design and in the other hand can explore a very important rout to employ PET in developing bioreactors for tis-sue engineering applications.

The numerical part includes the step 3 and 4 in which the numerical study has been performed for 3-D bone tissue growth in HFMB as a case study for large-scale tissue culture. In the step 3, the feasibility of utilizing newly proposed hHFMB for the growth of mesenchymal stem cells (MSCs) to form bone tissue was investigated using numerical simulations. To this aim, a mathematical model using a CFD code was developed to optimize the design and operation parameters of hHFMB for the growth of MSCs. The volume averaging method was used to formulate mass balance for the nutrients and the cells in the porous extracapillary space (ECS) of the hHFMB. The cell-scaffold construct in the ECS of the hollow fibres and membrane wall were treated as porous medium. Cell volume fraction dependent porosity, permeability, and diffusivity of mass were used in the model. The simulations allowed the simultaneous prediction of nutrient distribution and nutrient-dependent cell volume fraction. In addition, this model was used to study the effects of the operating and design parameters on the nutrient distribution and cell growth within the bioreactor. The modeling results demonstrated that the fluid dynamics within the ECS and transport properties and uptake rates in hHFMB were suf-ficient to support MSCs required for clinical-scale bone tissue growth in vitro and enabled to solve nutrition difficulties because of high cell density and scaffold size.

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In the step 4, the new dynamic and two-porous media model has been used for analyzing the nutrient-dependent MSCs growth in order to form the bone tissue in the HFMB. In the present model, hollow fibre scaffold within the bioreactor was treated as a porous domain. The domain consists of the porous lumen region available for fluid flow and the porous ECS region, filled with collagen gel containing cells, for growing tissue mass. Furthermore, the contributions of several design and process parameters, which enhance the performance of the bioreactor, were studied. In addition, the dynamic evaluation of cell growth, oxygen and glu-cose distributions were quantitatively analyzed. The obtained information can be used for bet-ter designing of the bioreactor, debet-termining of suitable operational conditions and scale up of the bioreactor for engineering of clinical-scale bone tissue.

In this study, the idea of using PET for visualization of steady and pulsatile flow perfusion within HFMBs which have proposed for the first time in this research project is quite original. Moreover, development of the scalable HFMB being compatible with MRI and PET, flat tissue culture chamber, combination of HFMB loop and a pneumatic system for generation of pulsatile wave forms and monitoring and control of the system has not reported in bioreactors used for tissue engineering applications by others. In addition, proposing hHFMB for clinical-scale bone tissue engineering application and modeling of nutrient distribution and MSC growth within hHFMB is quit original idea. Finally, development of a novel two-porous media model for mass transfer and MSC growth in a HFMB for tissue engineering of bone have been proposed and performed in this thesis.

Keywords: Hollow fibre membrane bioreactor; Tissue engineering; Large-scale tissue

con-struct; Positron emission tomography; Numerical modeling; Pulsatile flow; Bone tissue; Mes-enchymal stem cell

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ACKNOWLEDGMENTS

I would like to thank the Chemical Engineering Department of Universite de Sherbrooke for the opportunity that gave me to complete my PhD study during the project years. I am particularly grateful to my supervisors, Prof. Amin Behzadmehr, Prof. Gervais Soucy and Prof. Gerard Lachiver for accepting the supervisory role in the research pro-ject, and leading towards my thesis defense.

I would like also to express my very deeply thanks to Prof. Behzadmehr in respect of his helpful advises and discussions for the completion of modeling sections of this thesis. Many thanks are given to doctoral committee for their valuable suggestions and supports. I would like to mention my appreciation to Prof. Charles Doillon for his kind advices and discussions from the very beginning of my PhD research project in biological aspect of the project.

Also, I wish to express my special gratitude to Prof. Mercadier for providing me FLUENT software and computer in his laboratory at the Mechanical Engineering Department of Universite de Sherbrooke. Thanks are also expressed to laboratory Thermo 18 of Prof. Galanis and Mercadier.

I also acknowledge the "contribution" of my wife, Dr. Afra Hadjizadeh, who was my "teammate" during the project years, for her useful scientific discussions, suggestions from the very beginning of this project, especially in biological aspect of the project and also flat bioreactor chamber design, in respect of 3-D tissue and scaffold construct. The laboratory of the Bioingenierie et Biophysique of Universite de Sherbrooke is ac-knowledged in respect of financial support for the materials required for the experimental part of this research project.

Special thanks are expressed to Mr. Marc G. Couture and his friendly technical assis-tance, especially for his expertise in the drawing of my designs and machining the

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biore-actor chambers, oxygenator, and compliance chamber and other technical assistance. Thanks are also expressed to Mrs. Denis Turcotte, Serge Gagnon, and Alain Levesque for their technical assistance.

I would like to acknowledge the glass workshop of Chemistry Department of Universite de Sherbrooke for their assistance in construction of glass parts of the bioreactor system. The Civil Engineering Department is also acknowledged for providing place during the first part of the bioreactor system construction.

I would like also to acknowledge undergraduate laboratory (unit operation) of Chemical Engineering Department for providing place during second part of the bioreactor system construction and testing.

Many thank to Dr Jacques Rousseau, Prof. Roger Lecomte and technicians of CIMS (Centre d'imagerie moleculaire de Sherbrooke) at the Department of Nuclear Medicine and Radiobiology at the Faculty of Medicine for their support during PET scan experi-ments.

The personnel of the Chemical Engineering and Biotechnology Department of the Fac-ulty of Engineering, especially Mme Louise Chapdelaine and Mme France for their gen-eral support and Benoit Cote for their much appreciated technical assistance are also ap-preciated.

I wish to acknowledge MSRT of Iran for providing financial supports (scholarship) for me to continue my studies during four years and three month towards my Ph.D. degree. I also wish to thank Mme Sonia Morin, Prof. Michele Heitz, Mme Badaroudine and Prof. Gerard Lachiver, whose unforgettable helps made it possible for me to complete this study.

I also express my deep appreciation to my son, Sepehr, for his patience and comprehen-sion during these past five years.

I would also like to take the opportunity to thank my family for the support they have provided me with throughout my entire life, including these past five years.

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I take this opportunity to thank those many fine people (friends) that I met in Sherbrooke (Canada), who have made time enjoyable for me during my five years stay in here. My special thanks for my all teammates in the group of OPPUS and Biogenie and graduate students at the Department of Chemical Engineering for very kind socialization during my study at the Universite de Sherbrooke.

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LIST OF FIGURES

Figure 2.1: Classification of cell/tissue culture bioreactors based on method of operation 12

Figure 2.2: Schematic diagram of static culture vessels 13 Figure 2.3: (a) spinner flask bioreactor, (b) wavy wall bioreactor (WWB) (reproduced

from Bueno, E.M. et al., 2005) 14 Figure 2.4: Schematic diagram of rotating wall perfused vessels: (a) SLTV and (b)

HARV 16 Figure 2.5: Schematic diagram of a wave bioreactor 16

Figure 2.6: Concentric cylinder bioreactor (reproduced from Saini, S. et al., 2003) 17 Figure 2.7: Radial flow bioreactor (reproduced from Iwahori, T. et al., 2005) 18

Figure 2.8: Schematic diagram of a parallel plate bioreactor 19 Figure 2.9: Schematic diagram of a direct perfusion bioreactor 20 Figure 2.10: Micro-fluidic bioreactor (reproduced from Leclerc, E. et al., 2004) 21

Figure 2.11: Rotating shaft bioreactor (reproduced from Chen, H.C. et al., 2006) 22

Figure 2.12: Schematic diagram of a hollow fibre membrane bioreactor 23 Figure 2.13: Pulsatile flow bioreactors for tissue-engineered cardiovascular: (a) heart

valve and (b) vascular grafts (reproduced from Hoerstrup, S.P. et al., 2000 and

Niklason, L.E. et al., 1999) 24 Figure 2.14: Interacting phenomena in a typical perfusion bioreactor during tissue

growth (adapted from Coletti, F. et al., 2006) 37 Figure 3.1: Schematic diagram of the bioreactor system 43 Figure 3.2: Isometric view of the cell and tissue culture chamber (hollow fibre

biore-actor) 44 Figure 3.3: (a) Tissue culture chamber (Flat bioreactor) and (b) a sectioned view of the

flat bioreactor 44 Figure 3.4: Schematic representation of the assembly process of the different scaffold

designs that can be accommodated in the cell and tissue culture chamber. Assembling of: (a) the hollow-fibre scaffold, (b) channeled scaffold, (c) tubular

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Figure 3.5: Geometries used for CFD analyses for Designs 1 to 4 are illustrated. Designs 1 to 4 consist of inlet and outlet manifolds and hollow fibre scaffolds (shaded region). Designs 1, 3, 4: Rt = 3mm (/?,., inlet tube radius); Design 2: Rt =2mm; for

all Designs:Rs= 10mm (Rs, hollow-fibre scaffold radius) 48

Figure 3.6: Pulsatile flow generation system 54 Figure 3.7: The developed bioreactor 57 Figure 3.8: The wide range of pressure-flow conditions under steady flows with water at

37°C: (a) flow rates vs. transmural pressure, (b) transmural pressures vs. flow rate 58 Figure 3.9: Dynamic responses achieved for DO and pH control using the bioreactor

system: a) set points for DO were sequentially changed from 8 mg/1 to 7.5 mg/L, 7.7 mg/L, and 5.5 mg/1 using a PID method and b) set points for pH were sequentially changed from 7.4 to 7.18 and 7.4 using a PI method. Ml99 culture medium was used at a flow rate of 110 ml/min and 37 °C. Control was done by

regulating air, O2 and CO2 gas flow rates in the gas exchange unit 59 Figure 3.10: The effect of the inlet and outlet manifold geometries and Reynolds

numbers (60, 180, 300, 479) on the stream line distribution in cylindrical manifolds (Designs 1 and 2) and conical manifolds (Designs 3 and 4) at

dimension-less pressure drop parameter K= 1.0E+5 62 Figure 3.11: Dimensionless analysis of simulation results: (a) the relationship between

axial velocity (Vr/Vs) and radial distance (r/Rs ) along the fibre bundle

(hollow-fibre scaffold) for Designs 1 to 4 at Re=300 and K=1.0E+5, (b) pressure drop versus axial distance (z/2Rs ) along the central axis of the bioreactor for Design 1

at Re=300 and K=1.0E+5, (c) axial velocity versus radial distance for Design 4 at different Reynolds numbers and K=1.0E+5, (d) axial velocity versus radial distance for Design 2 at different Reynolds number and K=1.0E+5, (e) The non-uniformity index {</>) versus Reynolds number for Designs 1 to 4 at K = 1.0E+5, (f) The non-uniformity index versus dimensionless pressure drop parameter, K,

for Design 4 at different Reynolds numbers 63 Figure 3.12: (a) tubular adjustable cell/tissue culture chamber (Hollow fibre bioreactor),

(b) flat bioreactor, and (c) scaffold cartridge in the flat bioreactor 65 Figure 3.13: Records of pressure profiles using water at 37°C for a pressure varying

between 80 and 180 mmHg and mean flow rate of 150 ml/min: (a) a pressure wave similar to that of arterial blood pressure in the aorta with a frequency of 1

Hz and (b) a pressure wave with a frequency of 2 Hz 66 Figure 3.14: Record of a pressure and flow rate profiles with a frequency of 4 Hz: (a)

mean flow rate 200 ml/min and pressure between 80 and 180 mmHg and (b)

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Figure 3.15: Records of pressure and flow rate profiles using water at 37°C by reproducing a narrow and peaky wave form at a frequency of 1.5 Hz: (a) mean flow rate of 50 ml/min and pressure between 15 and 45 mmHg and (b) mean flow

rate of 180 ml/min and pressure between 90 and 120 mmHg 68 Figure 3.16: Acquired waveforms showing pressure and flow rate profiles using water at

37°C and frequency of 1.25Hz: (a) mean flow rate of 50 ml/min and pressure between 5 and 60 mmHg and (b) mean flow rate of 180 ml/min and pressure

between 80 and 130 mmHg 69 Figure 4.1: Sketch of the experimental setup incorporating the hollow membrane fibre

bioreactor 75 Figure 4.2: Sherbrooke's animal PET scanner and experimental setup 75

Figure 4.3: View of the activity profiles in bioreactor for the case of void ECS in which the injected 18FDG in the compliance chamber was pre- and well-mixed. The

image was acquired using 2.57 mCi tracers with flow rate of 70 ml/min in distilled water. Image (a) is coronal or side view, (b) is a transaxial view at Z/R=l, (c) is a sagittal or top view and (d) is a 3-D reconstruction of the activity

distribution 79 Figure 4.4: Normalized number of events vs. time: (a) in the dynamic and (b)

volum-etric scans for four different operating conditions. For each figure, the events number in each profile was normalized to the number of events at steady inlet flow (70 ml/min) in order to obtain a relative event number in each experiment at

each time point 81 Figure 4.5: PET image of steady inlet flow (70 ml/min) at the upstream face (Z/R=0.5)

of the fibre bundle: (a) reconstructed image at t=10 min, (b) contour plot of the normalized activity at t=10 min, and (c) normalized activity profiles at t=10, 15,

20, 25, and 30 minutes. The data are corrected for 18FDG decay 84

Figure 4.6: PET image of steady inlet flow (140 ml/min) at the upstream face (Z/R=0.5) of the fibre bundle: (a) reconstructed image at t=10 minutes, (b) contour plot of the normalized activity at t=10 minutes, and (c) normalized activity profiles at

t=10, 15, 20, 25, and 30 minutes. The data are corrected for 18FDG decay 85

Figure 4.7: PET image of pulsatile inlet flow (70 ml/min and 75 Beat/min) at the upstream face (Z/R=0.5) of the fibre bundle: (a) reconstructed image at t=10 min, (b) contour plot of the normalized activity at t=10 minutes, and (c) normalized activity profiles at t=10, 15, 20, 25, and 30 minutes. The data are corrected for

18FDG decay 86

Figure 4.8: PET image of dynamic scan: (a) and (b) for steady inlet flow (70 ml/min), (c) and (d) for pulsatile inlet flow (70 ml/min and 75 Beat/min), at t=30 minutes in the bed (fibre bundle, Z/R=2.6). Data has been normalized to bulk activity at

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Figure 4.9: Represents the activity profiles in the radial direction at four different axial positions. Data were normalized to bulk activity of steady flow with 70 ml/min

flow rate to obtain a relative activity concentration of tracer in each slice 90 Figure 5.1: (a) Drawing of a hollow fibre bioreactor with adjustable chamber of 0.5 cm

to 10 cm containing a scaffold; (b) Hollow fibre scaffold with effective fiber length of 5 cm simulated in the present work; (c) schematics of single hybrid hollow fibre: Lumen (Region I), Membrane (Region II), Porous ECS (Region III);

(d) the volume averaging 96 Figure 5.2: Magnified portion of the mesh and detail of the computational domain (a)

and identification of the boundaries (b) used in the simulation. There are 18000

elements in total for whole scaffold length (50 mm) 103 Figure 5.3: Grid independence test; the effect of mesh size on the radial oxygen

concentration (dimensionless) at ECS 106 Figure 5.4: Radial and axial normalized oxygen glucose concentration profiles for

num-erical modeling (in this work) compared to (a) mathematical modeling in a channeled porous scaffold for growing cardiac tissue (Radisic, M. et al., 2005),

(b-d) analytical solution for a HFMB(Heath, C. et al., 1987) 109 Figure 5.5: Flow streamline in the membrane and ECS (a) and velocity vectors in the

lumen (b).sa= 0.0028 I l l

Figure 5.6: Comparison of the contours plots of pressure, sa =0.0028 112

Figure 5.7: Comparison of the contours plots of pressure, sa = 0.344 112

Figure 5.8: Comparison of the contours plots of physical velocity (jum/ s) atsa = 0.0028.... 113

Figure 5.9: Comparison of the contours plots of physical velocity {fjml s) dXsa = 0.344 113

Figure 5.10: Comparison of the contours plots of dimensionless oxygen concentration in

lumen, membrane, and ECS. sa =0.0028 114

Figure 5.11 Comparison of the contours plots of dimensionless oxygen concentration in

lumen, membrane, and ECS. sa = 0.344 114

Figure 5.12: Radial concentration profiles of oxygen (a and b) and glucose (c and d) for

different dimensionless culture time at Z/L=0.1 and 0.9 117 Figure 5.13: Specific growth rate (dimensionless) (a) and oxygen uptake rate

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Figure 5.14: (a) Cell volume fraction over 15 dimensionless culture time for Z/L=0A

and (b) Radial variation of cell volume fraction at tltd = 15 for different Z/L 118

Figure 5.15: Radial variation of effective oxygen (a) and glucose (b) diffusion coeffi-cients in cell layer at tltd=\5 for different Z/L, and temporal variation of

effective oxygen (c) and glucose (d) diffusion coefficients in cell layer over 15

dimensionless culture time for Z/L= 0.1 118 Figure 5.16: Axial oxygen (a) and glucose (b) concentration profiles (dimensionless) in

hHFMB for t =15td at various radial positions 119 Figure 5.17: (a) Temporal variation of the average cell volume fraction; (b) average

oxygen and glucose concentrations; (c) difference in inlet and outlet average oxygen concentration (dimensionless); and (d) the difference in inlet and outlet

average glucose concentration (dimensionless) at different initial cell density 120 Figure 5.18: Temporal variation of the average: (a) cell volume fraction, (b) oxygen

concentration, (c) glucose concentration in cell layer for different inlet

concen-trations of oxygen and glucose 122 Figure 5.19: Temporal variation of the average cell volume fraction at different cell

death rates 123 Figure 5.20: Temporal variation of average cell volume fraction for different: (a)

mem-brane thickness and (b) ECS thickness. 124 Figure 5.21: Temporal variation of average cell volume fraction for various lumen Re

number 126 Figure 5.22: Average dimensionless concentration profiles of oxygen and glucose at

different initial cell volume fraction vs. axial Pe number in the lumen. The average dimensionless concentration profiles of oxygen at the cell layer (ECS) (a) and the outlet (b).The average dimensionless concentration profiles of glucose at

the cell layer (c) and the outlet (d) 126 Figure 5.23: Evaluation of consistency between hMSCs growth within a PET construct

in a perfusion system (Zhao, F. et al., 2005b) and the results obtained from model

(developed in this study) for hMSCs growth in an hHFMB 127 Figure 6.1: Diagram of the bioreactor and the simulation boundaries (a) and hexagonal

arrangement of the hollow fibres and gel matrix between the fibres (b) 134 Figure 6.2: Number of fibres,^ > for different inter-fibre spacing and outer fibre radius

(a), lumen region porosity with different inter-fibre spacing and inner fibre radius, R0 = 200fjm (b), and ECS region porosity (volume available for tissue

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growth) with different inter-fibre spacing and fibre outer radius (c), module

radiusi?m = 0.0 \m, fibre wall thickness Sm =50/jm,Lf =Lfw 141

Figure 6.3: Grid dependence test: Effect of grid distribution on the dimensionless

oxygen concentration profiles in radial direction (a), and in axial direction (b) 146 Figure 6.4: Model-data comparison for the growth of cells in extracapillary space (a)

and the outlet dimensionless mixing cup oxygen concentration (b) of the HFMB

vs. dimensionless time 149 Figure 6.5: Temporal evaluation of the average cell volume fraction s at different

membrane wall thickness, 8m 152

Figure 6.6: Temporal evaluation of the average cell volume fraction sat different

hindrance factorKm 152

Figure 6.7: Temporal evaluation of the average cell volume fraction s at different inner fibre radius i?,(a) and average cell volume fraction vs. R.fox several

dimension-less time (b) 154 Figure 6.8: Temporal evaluation of the average cell volume fractions at several

inter-fibre spacing (a) and average cell volume fraction vs. As for a different

dimen-sionless time (b) 155 Figure 6.9: Lumen inlet velocity profile (a), lumen velocity vectors for a portion of

computational domain (b), contour plots of dimensionless concentrations of oxygen (c) and glucose (d) in the lumen region, oxygen (e) and glucose (f) in the ECS region. The cell volume fractions in the ECS are presented for the case of

variable inlet velocity (g) and uniform inlet velocity (h) 157 Figure 6.10: Axial evaluation of the oxygen and glucose concentration fraction for

several fibre lengths 159 Figure 6.11: Axial pressure drop vs. inner radius of fibres (a), volume flow rate (b), and

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LIST OF TABLES

TABLE 4.1: A SUMMARY OF THE PERFUSION EXPERIMENTS CONDUCTED

IN THIS STUDY 77 TABLE 5.1: INITIAL AND BOUNDARY CONDITIONS FOR MOMENTUM

CONS-ERVATION EQUATIONS, MASS EQUATIONS, AND CELL BALANCE

CORRESPONDING TO THE FIGURE 5.2 104

TABLE 5.2: GRID DEPENDENCE TEST 106 TABLE 5.3: MODEL PARAMETERS AND VALUES USED IN THIS WORK WITH

THEIR TYPICAL RANGES AND REFERENCES 107 TABLE 6.1: MODEL PARAMETERS AND VALUES USED IN THIS STUDY WITH

THEIR TYPICAL RANGES AND REFERENCES 147 TABLE 6.2: MODEL PARAMETERS AND VALUES USED FOR MODEL

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LIST OF ABBREVIATIONS

CCD CFD CT DO EC ECM ECS EtO 18FDG FDG-PET FOV HARV HFMB hHEMB hMSC KCM LDV MAVA MSC MRI NMR PEP PEPT PES PET PET PvFB PIV PMM ROI RWPV SLTV TE TEC 2-D 3-D WWB Charge-Coupled Device Computational Fluid Dynamics Computed Tomography Dissolved Oxygen Endothelial Cell Extra-Cellular Matrix Extracapillary Space Ethylene Oxide Fluro-Deoxy-Glucose Fluro-Deoxy-Glucose-Positron-Emission-Tomography Field-Of-View

High Aspect Ratio Vessel

Hollow Fibre Membrane Bioreactor

Hybrid Hollow Fibre Membrane Bioreactor Human Mesenchymal Stem Cell

Krogh Cylinder Moddel Laser Doppler Velocimeters Macro Area Weighted Averaging Mesenchymal Stem Cell

Magnetic Resonance Imaging

Nuclear Magnetic Resonance spectroscopy Positron Emission Profiling

Positron Emission Particle Tracking Polyethersulfone

Polyethylene Terephthalate Positron Emission Tomography Radial Flow Bioreactor

Particle Imaging Velocimeters Porous Media Model

Region of Interest

Rotating Wall Perfused Bioreactor Slow Lateral Turning Vessel Tissue Engineering

Tissue Engineering Construct Two-Dimensional

Three-Dimensional Wavy-Wall Bioreactor

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CHAPTER 1 INTRODUCTION

1.1. Problem statements

Tissue engineering offers significant promise as a viable alternative to current clinical strate-gies for repair and replacement of the damaged tissue or organs by transplanting functional tissue engineering constructs (TECs) that are grown in vitro. Tissue engineering generally consists of five phases including cell harvesting, cell expansion, scaffold seeding, bioreactor culture and conditioning, and implantation. Each phase is still subject to research and there are no definitively established protocols yet. However, although promising results have been ob-tained in all phases, in general, the size of engineered tissues (e.g. bone,...) are presently in-sufficient for clinical application due to the lack of micro-vascular network found in the natu-ral tissue. This is because , relying on diffusion of the nutrient from outside and waste material from inside the scaffold would limit the size of engineered viable tissue to less than 5 mm thick (Das, D.B., 2007; Freed, L.E. and Vunjak-Novakovic G.., 1998) which is not as much valuable as large tissue mass in clinical practice. In recent years, there is a great need for clinical-scale TECs (Meinel, L. et al., 2004; Zhao, F. et al., 2006). However, despite a signifi-cant progress in tissue engineering techniques, the lack of clinical-scale engineered tissue for

in vivo implantation is still a major problem. To this aim, the study on large-scale engineered

tissue is the subject of an intensive research in recent years.

One of the important requirements for the development of three-dimensional (3-D) TECs is a bioreactor system that can support long-term tissue growth and promote uniform 3-D tissue formation for tissue replacement. Because in vitro development of 3-D and large-scale TEC is limited by nutrient and waste product transfer hindrance (Ellis, M.J. et al., 2007) using in vitro conventional cell culture techniques to produce 3-D thick TECs such as bone is not a promis-ing strategy (Dunn, J.C.Y. et al., 2006), because, in traditional 'porous block' scaffolds the formation of functional TEC hinders by nutrient limitations and waste product transport when the dimension of the scaffold and density of the cells increase within the scaffold (Ellis, M.J. et al., 2007). This results a non-homogeneous growth of cells that reduce functionality of the

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formed TEC from periphery to center of the scaffold. To overcome this problem, one possibil-ity is to employ a driving force on the culture media (fluid) that passes through the scaffold placed in a bioreactor (Ellis, MJ. et al., 2007). However, in this strategy the fluid shear stress increases due to an increase in the media flow rate (Cartmell, S.H. et al., 2003; Ellis, M.J. et al., 2007; Zhao, F. et al., 2006), which may affect the performance of the tissue production. For example, the high flow rate can damage the cells due to high shear stress and remove them from the scaffold surface (Ellis, M.J. et al., 2007). The above mentioned difficulties have been observed in high dense 3-D TECs in long-term culture period. This phenomenon can be explained by previously published studies reporting that the growth of cells is limited to 20-200 microns from the fluid-tissue interface in ex vivo (Kellner, K. et al., 2002). Regarding the above discussion, a better alternative to above mentioned systems can be the use of hollow fiber membrane bioreactor (HFMB) systems. In this strategy the obstacles caused by diffusion limitations or shear stress related problems on cells or tissue has been solved. In fact, this sys-tem can be an alternative equivalent for vascular network required in most cell-based ap-proaches for tissue regeneration.

To date, steady fluid flow has been used in flow perfusion bioreactors. However, as natural tissues such as bone and cartilage are subjected to dynamic loading (Jaasma, M.J. et al., 2008), to create a condition close to that of natural tissue the research has focused on the using pulsa-tile flow perfusion systems and the effects of pulsapulsa-tile flow on tissue growth have been stud-ied (Kofidis, T. et al., 2003a; Nguyen, D.T. et al., 2005; Sodian, R. et al., 2002). The results obtained from using pulsating flow bioreactors have revealed that the pulsatile action have helped to extend cell lifetime and to improve the performance of the bioreactor. Taken together, in order to improve in vitro nutrition of large-scale TEC and to overcome the prob-lem of engineered tissue size, development of advanced bioreactors are needed which is one of the objectives of this thesis (objective 1). This is because, advanced bioreactors are essen-tial for growing functional tissue and play important role in providing an optimized and con-trolled environment for 3-D tissue development.

In addition, in order to improve designing of the above mentioned bioreactors and tissue growth within them, efficient imaging techniques and theoretical modeling techniques are needed for monitoring and optimizing of the flow perfusion, biochemical parameters, and tis-sue growth during culture period in vitro.

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A major limitation in understanding HFMB performance for tissue culture lies in our inability to obtain direct information on conditions within the bioreactor and their effect on biochemis-try of the cells (Williams, S.N. et al., 1997). This is due to small size of hollow fibres (Das, D.B., 2007) that makes it too difficult directly measuring of the interests. While, to design or model a HFMB for large-scale tissue growth in vitro (e.g. bone and cartilage tissue engineer-ing) under steady or pulsatile flow conditions adequate knowledge on the fluid flow perfusion is required. Such knowledge is vital to the optimization of bioreactor culture conditions. In these bioreactors, nutrients and waste products distributions as well as cell growth in the ECS and shear stress in the lumen are affected by the overall and local fluid flow rates. The latter becomes important when only cells are cultured on the lumenal surface of the fibres. Medical imaging technique like PET can be helpful for bioreactor design and optimization, which can be further a supplement to other techniques used for visualizing flow perfusion in the porous scaffolds placed in a perfusion bioreactor.

PET allows qualitatively and quantitatively measuring the spatial distribution of a radioactive tracer by detecting its activity within an opaque and porous object. This technique is widely used in the field of nuclear medicine (Bentourkia, M., 2003; Selberg, O. et al., 2002). But, it has never been employed for imaging inside HFMBs as a complex opaque system to visualize flow perfusion, mass transport and metabolic activity of growing cells under steady and pulsa-tile flow.

In addition, computational modeling can aid in a more structured approach. These mathemati-cal frame works can help to establish quantitative relationships in order to direct experiments and reduce number of them. Moreover, these models would aid to estimate parameters and variables which can not be measured directly by experimental methods (Mehta, K. et al., 2006). These include examining suggested hypotheses and relationship between observations, addressing the experimental issues and understanding of the system itself, thereby opening up new routes for research. Optimal flow and mass transport conditions as well as nutrient pendent cell growth in the bioreactor can be supported by simulation methods which need de-veloping efficient mathematical frame works. Computational modeling also enables to predict the development of functional tissue engineered constructs as a function of the physical, bio-chemical, and biological parameters provided in advanced bioreactor environment. Mass and momentum transport in perfusion bioreactors for cell/tissue growth has been studied in

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vari-ous contexts using mathematical modeling and CFD models (Williams, K.A. et al., 2002). These models have been developed in order to characterize the microenvironment for bioreac-tor designs (Abdullah, N.S. et al., 2006; Pathi, P. et al., 2005; Zhao, F. et al., 2007). In connec-tion with the aforemenconnec-tioned discussions, the scope and objectives of the present thesis will be given in the next section.

1.2. The scope of this thesis

The overall objective of this thesis is that of developing an advanced bioreactor system to both large-scale and long-term tissue culture within a 3-D environment. In order to achieve the mentioned aims, the following three specific objectives have been considered for this study:

1. Design, development, construction, and hydrodynamic evaluation of pulsatile flow perfusion bioreactor system for use with hollow fibre membrane scaffolds.

2. Visualization of fluid flow perfusion in the hollow fibre membrane bioreactor using a biomedical noninvasive imaging technique.

3. Developing two mathematical models for nutrient transport and tissue growth within a hybrid HFMB (hHFMB) and HFMB with the aid of CFD.

Objective 1: Design, construction, and hydrodynamic evaluation of a pulsatile flow perf-usion bioreactor system for use with hollow fibre membrane scaffolds.

Currently available bioreactors have serious limitations in terms of supporting large-scale and long-term tissue culture. A major challenge is to design a tissue culture bioreactor system that efficiently supports tissue-based constructs.

In this study, a continuous pulsatile flow bioreactor system is designed and constructed to meet the general and specific requirements of developing large-scale tissue constructs. An ideal bioreactor should meet the both biochemical and biomechanical controls over the devel-oping tissue. In order to control biochemical parameters such as pH, pC>2, pCC>2, a gas ex-change unit is designed to use the system without standard incubator. To control

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biomechani-cal parameters in term of hemodynamic flow, an independent pulsatile flow generation system is designed and constructed. As it is necessary to image flow perfusion and growing tissue in the bioreactor during culture period without disturbing the tissue construct, the system is de-signed to be compatible with medical imaging techniques such as MRI and PET available for this research project. Generally, facilities for medical imaging (i.e. medical imaging machines) are centralized in medical centers due to their required utilities and applications. Therefore, the system is designed to be compact and portable in order to port to medical imaging center without shutting down long-term experiments. One of the important factors on bioreactor de-sign is considering a sterilization method. Generally, most of commercially available bioreac-tors and even bioreacbioreac-tors that have been used for research propose to use ethylene oxide (EtO) for serializations. In this work, it is considered that the bioreactor to be designed compatible with autoclave which has no any toxic effect after sterilization. In order to monitor and control the parameters governing the system, Lab VIEW software and a home-made program is used. Important unit of a bioreactor is cell/tissue culture chamber. Two different tissue culture chamber are designed and constructed which offer the following advantages: 1) capable to ac-commodate hollow fibre scaffold, 2) scalable in size, 3) MRI and PET compatible, 4) fully visible, 5) enable the user to easily fix the scaffold, 7) compatible with CCD camera and dif-ferent microscopes, 8) reusable, 9) stable in long-term culture period.

Other aspect which is considered in this design is that the bioreactor loop should be separable from the system in order to sterilize in the standard autoclave and provide cell seeding under standard laminar flow hood and then re-attach to the system. Designing the inlet and outlet manifold of the bioreactor is performed by CFD simulation of the flow distribution to achieve high efficiencies in the uniformity of flow perfusion.

In the current study, a collapsible tube and a pneumatic system (pulsatile flow generator) are used to generate pulsatile flow. This method has advantages over other systems for creating flow pulsation in that there are no needs to sterilization. In addition, perfusion fluid does not come into contact with oil or other foreign matter. Moreover, it can work independently with perfusion pump resulting independence of flow rate and pulsatile flow. As mentioned earlier, past studies (Kofidis, T. et al., 2003a; Nguyen, D.T. et al., 2005) reported in literature have demonstrated that the pulsatile flow is able to improve performance of the bioreactors for

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tis-sue engineering applications. As the author's awareness, no pulsatile flow bioreactor study with hollow fibre membrane scaffold has been reported so far. The detailed design, construc-tion, and hydrodynamic evaluation of the bioreactor are outlined in Chapter 3.

Objective 2: Visualization of fluid flow perfusion in the hollow fibre membrane biore-actor using a biomedical noninvasive imaging technique

There has been significant progress in growing tissue in vitro. However, the scale up of such technology to clinical setting remains a significant problem. This includes identifying the op-timal flow parameter, scaffold, and bioreactor geometry for the generation of clinical-scale tissue. To do this, there is a need for methods that can be used to assess the influence of perfu-sion and nutrient diffuperfu-sion on bioreactor performance in terms of cell growth and distribution without disturbing growing tissue within opaque bioreactor.

In this study, a noninvasive technique known as Positron Emission Tomography (PET) is pro-posed for the first time and used to image steady and pulsatile flow patterns in a HFMB for tissue engineering applications. To this aim, a small animal PET scanner is used. The closed loop bioreactor system, mentioned in objective 1, is employed in this study. Several dynamic scans were performed and then followed by volumetric scans of whole bioreactor, which then were used to obtain dynamic and volumetric 3-D images. These images were used for qualita-tively and quantitaqualita-tively analyzing of flow perfusion in the bioreactor. The detailed experi-mental approach, analysis, and results are outlined in Chapter 4 of this thesis.

Objective 3: Developing mathematical models for nutrient transport and tissue growth within hHFMB and HFMB with the aid of computational fluid dynamics (CFD).

Producing clinical-scale engineered bone tissue using traditional cell culture methods in vitro is still a challenging issue, due to nutrient and oxygen transfer limitations. To address this obstacle, the use of HFMBs, mimicking the capillary networks that exist in natural tissues, have been previously proposed for 3-D tissue growth (Chesnick, I.E. et al., 2007; Ye, H. et al., 2007). However, for an efficient design of above mentioned bioreactors, the relationships between the growth of cells and their environments should be determined by using efficient numerical methods. To this aim, In the second part of the thesis (Chapter 5, 6), as a case study,

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I have focused on the mathematical modeling of large-scale in vitro expansion of Mesenchy-mal Stem Cells (MSC) for bone tissue in HFM-based bioreactors.

In Chapter 5 of this study, the use of a hollow fibre having a porous matrix around that which is called hybrid hollow fibre membrane bioreactor(hHFMB) is proposed to grow MSCs for bone tissue formation that has been used for bioartificial liver assist devices (Hoque, M.E. et al., 2007; Mareels, G. et al., 2006) . This system both plays the role of artificial vessels to support fluid flow and cell attachment by creating porous ECS while protects cells from shear stress. The feasibility of utilizing the above proposed hHFMB for the growth of MSCs to form bone tissue is investigated using numerical simulations. To this aim, a mathematical model by using a computational fluid dynamics (CFD) code is developed to optimize the design and op-erating parameters affecting the growth of MSCs in hHFMB for a long-term culture period. The volume averaging method is used to formulate mass balance for the nutrients and the cells in the porous ECS of the hHFMB in which effective diffusivity of nutrients, porosity, and permeability of porous ECS spatially and temporally change with growing cells. The cell-scaffold construct in the ECS of the hollow fibres and membrane wall are treated as porous medium. The detailed modeling procedure is outlined in Chapter 5 of the thesis.

As presented in Chapter 6 of the present study, a novel two porous media model to study the MSCs growth in a HFMB for bone tissue engineering applications is also developed. In this bioreactor, cells are grown in collagen matrix filled in the ECS of the bioreactor. These cells then proliferate in the gel matrix during a long-term culture period. A global and dynamic model is developed to simulate nutrient dependent cell growth patterns, as well as to account for nutrient transport and uptake rates in the bioreactor by determining of effective diffusion coefficients and uptake rates for tissue formation in terms of temporal and spatial restrictions. This study would advance our understanding of parameters that affect cell growth in large-scale and long-term bone tissue formation in vitro in HFMB. Moreover, the model will lead to develop these bioreactors with optimal design and operating parameters for growing engi-neered bone tissue. Furthermore, the developed model could enable to provide suggestions in order to improve long-term experimental design by identifying the effective parameters which need to be taking into account. This work indirectly can be useful for monitoring the devel-opment of the engineered bone tissue construct in HFMB. The detailed modeling approach is outlined in Chapter 6 of the dissertation.

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In summary, this thesis consists of seven chapters. The project-related literature is reviewed in Chapter 2, including several bioreactor designs, techniques used for bioreactor imaging, bio-medical background of small animal PET, and reported tissue growth improvement by pulsa-tile flow, CFD and its role in tissue engineering and bioreactor analysis for tissue growth and mathematical modeling of HFMBs. A pulsatile flow bioreactor system is designed and con-structed. Hydrodynamic and control evaluations are performed in the study, and detailed design, development, CFD analysis of the bioreactor and results has been presented in Chapter 3.

The PET method for imaging of the steady and pulsatile flow perfusion in a hollow fibre bio-reactor has been explained in Chapter 4, including the designing setup, conducted experi-ments, and results.

In the second part of the thesis, a novel numerical dynamic modeling of a hybrid hollow fibre bioreactor for bone tissue engineering application from MSC has been presented in Chapter 5, including theoretical background, developing mathematical model, validation, and results. A novel global modeling (two-porous media model) of the hollow fibre bioreactor with MSC in the ECS of the bioreactor is developed in Chapter 6 including theoretical background, mathematical modeling, validation, and results.

The conclusions of this study are given in the last chapter of the thesis, Chapter 7, as well as some recommendations for future research studies.

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CHAPTER 2 BIOREACTORS FOR TISSUE ENGINEERING: LITERATURE

REVIEW AND BACKGROUND

2.1. Overview

One of the most challengeable problems in tissue engineering is in vitro formation of clinical-scale tissues using traditional tissue engineering protocols for which the supply of nutrients controls the success or failure of tissue formation. Particularly, this is more effective for a large-scale engineered tissue due to high cell density in most tissues within a large-scale scaf-fold in vitro. To address this issue, one of the most promising approaches in tissue engineering to date can be the use of cell/tissue culture systems (i.e., bioreactors) that are able to supply sufficient nutrient to the cells with continuous waste removal. It is expected that this approach allows the culturing of cells in higher density and developing tissue/organs in vitro. Among the continuous bioreactor systems, perfusions systems play an important role for continuous culture and are widely used to culture cells in tissue engineering. Advances in bioreactor de-velopment are however necessary in order to attain the ultimate goal of a functional engi-neered tissue. An advanced bioreactor system for clinical-scale tissue engineering needs more complex design such as automated bioprocess, monitoring environment and tissue growth visualization, having an efficient cell growth conditions, facilitate their broad clinical use. To this aim, this chapter focuses on the reviewing and describing conventional and new ap-proaches on the bioreactor systems, bioreactor visualization methods, computational fluid dy-namics, and mathematical methods for bioreactor design.

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2.2. Introduction

Bioreactors have been designed to provide biological and/or biochemical processes develop-ment by monitoring and controlling of environdevelop-mental and operating conditions in vitro (e.g. pH, temperature, pressure, nutrient and growth factors supply and waste removal) close to that of in vivo. In addition, they should provide control, automation and reproducibility which are three important parameters for transferring such technology to large-scale applications (Mar-tin, I. et al., 2004). Specifically bioreactors are utilized for different purposes including: (1) cell proliferation on small-scale and large-scale, (2) production of 3-D tissue constructs in

vi-tro using previously isolated and proliferated cells and (3) as devices that play the role of

di-rect supporting of organ (i.e. artificial organs) (Portner, R. et al., 2005; Shachar, M. et al., 2003). In this brief review, first the key technical challenges are identified and then an over-view of developed culture vessels and bioreactors employed for tissue engineering is pre-sented. In the second part of this chapter noninvasive techniques which are used for bioreactor development in tissue engineering are reviewed and then a brief review of mathematical mod-eling of HFMBs for tissue engineering applications is presented.

2.3. Important requirements and critical parameters in bioreactor design

for tissue engineering (TE)

The design and development of bioreactors for tissue engineering applications is an important engineering problem. Therefore, it needs understanding of engineering and scientific back-grounds in order to develop mechanically and physiologically an appropriate environment for the growth of specific tissue. Each design tends to be specific to a given application. Ideally, bioreactors used to support tissue growth must provide the folio wings:

1) Maintenance of satisfactory biochemical conditions (e.g., pH, pC>2, concentrations of nutrients and growth factors, etc.) to support cell proliferation and/or differentiation into functional tissues.

2) Sufficient nutrient transport to the developing tissue. 3) Efficient waste removal from the developing tissue.

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5) Easy manipulation of the scaffolds under strict sterile conditions.

6) Utilization of existing techniques for monitoring cell/tissue growth such as non-invasive medical imaging techniques

7) Mechanical forces such as compression and expansion for growing tissue as well as provide homodynamic forces such as shear stress and pressure as found in vivo. 8) A high degree of reproducibility, control and automation.

9) Control flow perfusion both in steady or pulsatile conditions. 10) Laminar fluid flow or reduce turbulence in the fluid flow. 11) A low volume capacity to reduce cost of experiments. 12) An efficient use of growth factors.

13) Avoids the accumulation of the metabolites.

14) Materials which are used for the fabrication of the bioreactor should be sterilizable and also compatible with tissue/cell.

To meet these requirements and advance the utility of tissue engineering, improved tissue cul-ture systems and their associated processes that can support long-term tissue growth and pro-mote uniform 3-D tissue formation are needed (Martin, I. et al., 2004; Sodian, R. et al., 2002). To achieve efficient tissue mass growth with an optimal nutrient supply and waste removal, perfusion techniques along with the identification of the optimal culture conditions such as pressures, flow rates, and types of flow regimes are essential. Few commercial bioreactor sys-tems allow flow perfusion and control over all these operating conditions. In the next section, a brief classification of current bioreactor designs is presented.

2.4. Classification of bioreactors based on method of operation

Different ways can be applied to classify the bioreactors including specific applications, ge-ometry, and method of operation. In this chapter, a short review of current bioreactor designs in cell/tissue engineering is presented in Figure 2.1. In the presentation, bioreactors are classi-fied based on method of operation including static, mixed and perfusion bioreactors.

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Bioreactors for tissue engineering

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Perfusion systems

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Mixed systems f Static systems J U T-Flasks Dishes Plates Roller bottles Spinner flask Rotating wall Wave Concentric cylinder Direct perfusion Rotary shaft Hollow fibre membrane I—f Micro-fluidic l_| Radial flow |—\ Parallel plate I—I Pulsatile flow

Figure 2.1: Classification of cell/tissue culture bioreactors based on method of operation

2.4.1. Static culture systems

Static culture systems are first and traditional bioreactors (Niklason, L.E. et al., 1999) which present a simple way to culture cell/ tissue in vitro including Petri dishes, bags, T-flasks, and roller bottles (Fig. 2.2). These devices consist of a compartment where nutrient and growth factors exist in the medium diffuse from outside of the scaffold to the cells and waste products from inside to the medium. There are many problems using these kind of culture system such as maldistribution of cells over the construct, high cell death rate due to lack of nutrients and oxygen and insufficient extracellular matrix (ECM) production (Ishaug, S.L. et al., 1997; Mar-tin, I. et al., 2004). In these systems penetration depth is not exceed to maximum 240 urn due to mass transfer limitations while the size of engineered tissue constructs is in order of mm. Therefore, external mass transfer limitations should be reduced by culturing constructs in the mixed or perfusion culture systems (Martin, I. et al., 2004).

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12 well plate Roller bottle Figure 2.2: Schematic diagram of static culture vessels

2.4.2. Mixed systems

Second type of the bioreactors presents an efficient method of mixing in order to reduce the stagnant boundary layer (or external mass transfer limitation) in the periphery of the cells and scaffolds which exist in the static culture systems. In mixed systems, a high degree of homo-geneity in term of nutrients and growth factors is provided for growing cell/ tissue. The design of such bioreactors makes it possible to employ fluid perfusion in addition to mixing if cells immobilized onto scaffold placed in the bioreactor. Various mixed bioreactors, for this aim, have been designed and used in tissue engineering including:

Spinner flask bioreactors

Spinner flask bioreactors (Fig. 2.3a) consist of a vessel made of plastic or glass, and a mag-netic stir bar or impeller, and two side arms for medium supply or removal and adding cells, and supplying gas. Using spinner flasks bioreactors have shown that the efficiency of cell seeding and viability of cells are improved when compared to static culture vessels. During cell seeding process, because of fluid convection suspended cells are transported into the pores of suspended scaffold. The efficiency of chondrocyte seeding onto polyglycolic acid scaffolds in spinner flasks has been reported to reach 100% within 1 day (Vunjak-Novakovic, G. et al., 1998). In these bioreactors, scaffolds in the form of porous blocks are suspended within the bioreactor by using needles. This kind of bioreactors has been used for cultivation of variety of tissue constructs including bone, cartilage, MSCs, microcarier suspension (Sucosky, P. et al., 2004; Vunjak-Novakovic, G. et al., 1999). Major drawback of using spin-ner flasks bioreactors is external and internal diffusions limitations and non-uniform tissue

Le développement et la modélisation numérique d'un bioréacteur pour l'ingénierie des tissus de grande masse

SHERBROOKJE

DEVELOPMENT AND NUMERICAL MODELING OF A BIOREACTOR

SYSTEM FOR THE ENGINEERING OF LARGE-SCALE TISSUE

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Composition du jury /Doctoral Committee:

Prof. Charles Doillon

Prof. Amin Behzadmehr

Prof. Bernard Marcos

Prof. Yves Mercadier

Prof. Gerard Lachiver

In memory of my loved and dearly-missed father and teacher

This work is dedicated to those I love dearly:

RESUME

ABSTRACT

ACKNOWLEDGMENTS

TABLE OF CONTENTS

LIST OF FIGURES

LIST OF TABLES

LIST OF ABBREVIATIONS

CHAPTER 1

INTRODUCTION

1.1. Problem statements

1.2. The scope of this thesis

CHAPTER 2

BIOREACTORS FOR TISSUE ENGINEERING: LITERATURE

REVIEW AND BACKGROUND

2.1. Overview

2.2. Introduction

2.3. Important requirements and critical parameters in bioreactor design

for tissue engineering (TE)

2.4. Classification of bioreactors based on method of operation

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