Applied Science, Faculty of
Engineering, School of (Okanagan)
DSpace
UBCO
Bhullar, Arpin Singh
Master of Applied Science - MASc
University of British Columbia
Valvular heart disease is a cardiovascular condition characterized by damage to any number of the four valves within the human heart. This damage can be defined as stenosis or incompetence based on a valveâs inability to completely open or close, respectively. Stenosis and incompetence are caused by calcification and the thickening of valvular tissue, and often manifest as angina, ultimately leading to congestive heart failure if left untreated. Mechanical heart valves (MHV) can be surgically implanted to replace a failing valve but require the prescription of life-long anticoagulant medication such as Warfarin to mitigate blood clotting issues. The thrombogenicity of MHVs is related to their poor hemodynamics, believed to be caused by the high shear stresses they induce. Some identified regions of high shear stress are the hinges, the trailing edges of the leaflet tips, and velocity jets. These problems have plagued the most common form of MHVs, bileaflet mechanical heart valves (BMHVs), for decades as their designs have stagnated. This research was conducted to design and validate a BMHV that would have similar hemodynamic properties to native heart valves, enabling implantation without reliance on post-operative anti-coagulants. To accomplish this, we identified the major design limitations of BMHVs, such as the gold standard St. Jude Medical Regent BMHV, as well as the believed contributing factors to their thrombogenicity. We developed a novel design architecture for BMHVs that addressed these limitations, and from it, the Okanagan Valve (OKV). A numerical analysis was performed on modifications of interest incorporated in the OKV to determine their effect on heart valve hemodynamics. In-vitro methods were utilized to experimentally observe the performance of the OKV. Our experimental assessment of the OKV found the maximum regional backflow velocity and closing volume to be 46 m/s and ~5 cc, respectively, values closer to those of a bioprosthetic valve rather than a BMHV. The results of our in-vitro assessment confirmed the results of the numerical analysis, showing the potential of the OKV to accomplish the design goal. Ultimately, this research is intended to be a proof of concept for the OKV, which may one day become a BMHV that can be implanted indefinitely without the need for anticoagulants.
Okanagan Valve: The Next Generation of Mechanical Heart Valves by Arpin Singh Bhullar B.ASc., The University of British Columbia, 2017 A THESIS SUBMITTED IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF MASTER OF APPLIED SCIENCE in THE COLLEGE OF GRADUATE STUDIES (Mechanical Engineering) THE UNIVERSITY OF BRITISH COLUMBIA (Okanagan) April 2020 © Arpin Singh Bhullar, 2020 ii The following individuals certify that they have read, and recommend to the College of Graduate Studies for acceptance, a thesis/dissertation entitled: Okanagan Valve: The Next Generation of Mechanical Heart Valves submitted by Arpin Singh Bhullar in partial fulfillment of the requirements of the degree of Master of Applied Science in Mechanical Engineering Examining Committee: Dr. Hadi Mohammadi, Mechanical Engineering Supervisor Dr. Ray Taheri, Mechanical Engineering Supervisory Committee Member Dr. Bahman Naser, Civil Engineering Supervisory Committee Member Dr. Liwei Wang, Electrical Engineering University Examiner iii Abstract Valvular heart disease is a cardiovascular condition characterized by damage to any number of the four valves within the human heart. This damage can be defined as stenosis or incompetence based on a valveâs inability to completely open or close, respectively. Stenosis and incompetence are caused by calcification and the thickening of valvular tissue, and often manifest as angina, ultimately leading to congestive heart failure if left untreated. Mechanical heart valves (MHV) can be surgically implanted to replace a failing valve but require the prescription of life-long anticoagulant medication such as Warfarin to mitigate blood clotting issues. The thrombogenicity of MHVs is related to their poor hemodynamics, believed to be caused by the high shear stresses they induce. Some identified regions of high shear stress are the hinges, the trailing edges of the leaflet tips, and velocity jets. These problems have plagued the most common form of MHVs, bileaflet mechanical heart valves (BMHVs), for decades as their designs have stagnated. This research was conducted to design and validate a BMHV that would have similar hemodynamic properties to native heart valves, enabling implantation without reliance on post-operative anti-coagulants. To accomplish this, we identified the major design limitations of BMHVs, such as the gold standard St. Jude Medical Regent BMHV, as well as the believed contributing factors to their thrombogenicity. We developed a novel design architecture for BMHVs that addressed these limitations, and from it, the Okanagan Valve (OKV). A numerical analysis was performed on modifications of interest incorporated in the OKV to determine their effect on heart valve hemodynamics. In-vitro methods were utilized to experimentally observe the performance of the OKV. Our experimental assessment of the OKV found the maximum regional backflow velocity and closing volume to be 46 m/s and ~5 cc, respectively, values closer to those of a bioprosthetic valve rather than a BMHV. The results of our in-vitro assessment confirmed the results of the numerical analysis, showing the potential of the OKV to accomplish the design goal. Ultimately, this research is intended to be a proof of concept for the OKV, which may one day become a BMHV that can be implanted indefinitely without the need for anticoagulants. iv Lay Summary Valvular heart disease is a heart condition that is defined by damage to any of the four valves within the human heart. This damage is categorized by either insufficient blood flowing through or excessive blood flowing backwards through the valve. These conditions are caused by calcium growth and thickening heart valve tissue. The most common symptom is chest pain, and the condition can be fatal if untreated. Mechanical heart valves (MHV) can be surgically implanted to replace a failing valve but require life-long consumption of blood thinners, such as Warfarin, to prevent blood clotting issues. This research was conducted to design and validate a bileaflet MHV (BMHV) that would perform like natural heart valves, enabling implantation without the need for post-operative blood thinners. Our tests with a heart simulator corroborated the results of our numerical study, showing potential for medication-free implantation of the Okanagan Valve. v Preface This thesis was prepared under the supervision and guidance of Dr. Hadi Mohammadi at the UBC Okanagan Campus. Dr. Mohammadi provided the research topic of designing a radical mechanical heart valve prosthesis that leveraged novel design elements to mitigate anticoagulant therapy and improve patient quality of life. All research presented henceforth was conducted at the Heart Valve Performance Laboratory at the University of British Columbia Okanagan Campus and at ViVitro Systems Inc. in Victoria, BC. Lawrence Scotten (ViVitro Systems Inc.), provided expertise and facilities for conducting in-vitro and steady-state tests of Okanagan Valve prototypes. Sections of Chapters 3, 4, and 5 were presented by the author at the 34th Annual Conference of the Canadian Biomaterials Society. A poster and an accompanying white paper were prepared for the conference. Content from Chapters 3, 4, and 5 have been assimilated into a paper that was submitted to the Journal of Medical Engineering & Physics for publication. A provisional US patent was filed and approved for the intellectual property contained within Chapter 3. Copyrighted images were reprinted with permission from multiple sources in Chapters 1, 2, and 4. Seven such images appear in Chapter 1, one image in Chapter 2, and two images in Chapter 4. vi Table of Contents Abstract .......................................................................................................................... iii Lay Summary ...................................................................................................................iv Preface ............................................................................................................................ v List of Figures ................................................................................................................... x List of Abbreviations ....................................................................................................... xv Acknowledgments ........................................................................................................ xvii Dedication ................................................................................................................... xviii Introduction and Background .......................................................................... 1 1.1 Native Heart Valves ................................................................................................ 2 1.2 Valvular Heart Disease ............................................................................................ 4 1.2.1 Stenosis and Incompetence ............................................................................ 5 1.3 Non-Mechanical Heart Valve Protheses ................................................................. 6 1.3.1 Bioprosthetic Heart Valves .............................................................................. 7 1.3.2 Transcatheter Heart Valves ............................................................................. 8 1.4 Mechanical Heart Valves ........................................................................................ 9 1.4.1 Caged Ball MHVs ........................................................................................... 10 1.4.2 Caged Disc MHVs ........................................................................................... 11 1.4.3 Tilting Disc MHVs ........................................................................................... 11 1.4.4 Bileaflet MHVs ............................................................................................... 11 1.5 Thesis Outline ....................................................................................................... 12 Design Limitations of Bileaflet Mechanical Heart Valves ................................ 13 2.1 Hemodynamics ..................................................................................................... 13 vii 2.1.1 Shear Stresses ............................................................................................... 14 2.1.2 Flow Stagnation ............................................................................................. 15 2.1.3 Regurgitation ................................................................................................. 15 2.1.4 Effective Orifice Area..................................................................................... 16 2.1.5 Cavitation ...................................................................................................... 18 2.2 Durability .............................................................................................................. 19 2.3 Material Biocompatibility ..................................................................................... 19 2.3.1 Valve Structure .............................................................................................. 20 2.3.2 Suture Ring .................................................................................................... 20 Design of the Okanagan Valve ....................................................................... 22 3.1 Leaflet Design ....................................................................................................... 23 3.2 Housing Design ..................................................................................................... 26 3.3 Hinge Design ......................................................................................................... 29 3.4 Critical Design Elements ....................................................................................... 30 3.4.1 Central Blood Flow ........................................................................................ 30 3.4.2 Effective Orifice Area..................................................................................... 31 3.4.3 Hinge Washing............................................................................................... 31 3.4.4 Surface Area .................................................................................................. 32 3.4.5 Shear Stress ................................................................................................... 33 3.4.6 Closing Volume .............................................................................................. 34 Methods ....................................................................................................... 35 4.1 Flow Simulation .................................................................................................... 35 4.2 Numerical Modeling ............................................................................................. 37 4.2.1 Modelling Setup ............................................................................................ 41 viii 4.3 Okanagan Valve Prototypes ................................................................................. 41 4.4 In-Vitro Mounting Gaskets ................................................................................... 46 4.5 In-Vitro Steady-State Leakage Testing .................................................................. 51 4.6 In-Vitro Pulse Duplicator Testing .......................................................................... 54 Results .......................................................................................................... 59 5.1 Numerical Results ................................................................................................. 59 5.2 In-Vitro Steady-State Leakage Results .................................................................. 61 5.3 In-Vitro Pulse Duplicator Results .......................................................................... 62 Conclusions and Future Works ...................................................................... 65 6.1 Discussion ............................................................................................................. 65 6.2 Limitations ............................................................................................................ 67 6.2.1 Prototype Manufacturing .............................................................................. 67 6.2.2 Prototype Material ........................................................................................ 67 6.2.3 Hinge Geometry Reproduction ..................................................................... 67 6.2.4 Minimal In-Vitro Testing Ability .................................................................... 68 6.3 Future Works ........................................................................................................ 68 6.3.1 Housing Inlet/Outlet Optimization................................................................ 68 6.3.2 Optimizing Ovality ......................................................................................... 68 6.3.3 Hinge Design .................................................................................................. 68 6.3.4 Leaflet Curvature Optimization ..................................................................... 69 6.3.5 Rapid Prototype Manufacturing and Local In-Vitro Analysis ........................ 69 6.3.6 Addressing Endocarditis ................................................................................ 69 6.3.7 Pediatric OKV ................................................................................................. 69 Bibliography .................................................................................................................. 71 ix Appendices .................................................................................................................... 77 Appendix A: FFF Prototypes ............................................................................................ 77 Appendix B: Flow Simulation .......................................................................................... 81 B.1 Side Views of Both Designs ............................................................................... 81 B.2 Angled Views of Both Designs .......................................................................... 83 B.3 Views of the Simulated Aortic Root .................................................................. 86 Appendix C: Okanagan Valve Renders ............................................................................ 87 C.1 Side view ........................................................................................................... 87 C.2 Front view ......................................................................................................... 88 C.3 Top view ............................................................................................................ 89 C.4 Bottom view ...................................................................................................... 90 C.5 Isometric view ................................................................................................... 91 Appendix D: Engineering Drawings ................................................................................. 92 D.1 Leaflet Drawing ................................................................................................. 93 D.2 Housing Drawing ............................................................................................... 94 x List of Figures Figure 1.1: (A) Heart location in the human body (B) Labelled cross-sectional view of the human heart. Reproduced with permission from National Heart, Lung, and Blood Institute; National Institutes of Health; U.S. Department of Health and Human Services. [2] .................................................................... 2 Figure 1.2: A healthy aortic valve in the open and closed positions. Modified with permission from Jill Rhead under a Creative Commons CC BY-NC-ND 4.0 license. [3] ........................................................................................................ 3 Figure 1.3: The anatomy and major components of the mitral valve. Reproduced with permission from [4], Copyright Massachusetts Medical Society. ............ 4 Figure 1.4: (A) Healthy Aortic Valve (B) Severely Calcified Aortic Valve. Adapted by permission from Springer Nature Customer Service Centre GmbH: Springer Nature, Nature Reviews Disease Primers, [Calcific aortic stenosis], B. R. Lindman et al, 2016. [8] ............................................................ 6 Figure 1.5: (A) Stented Medtronic Hancock II BHV (B) Stentless Edwards Prima Plus BHV. Adapted by permission from BMJ Publishing Group Limited. [Transcatheter heart valve implantation for failing surgical bioprostheses: technical considerations and evidence for valve-in-valve procedures], D. Mylotte et al, 99, 960-967, 2013. [12] ................................... 8 Figure 1.6: (A) Edwards-Sapien Valve (B) Medtronic CoreValve. Reproduced with permission under a Creative Commons CC BY-NC-ND 4.0 license. "Transcatheter aortic valve replacement: design, clinical application, and future challenges." by J. K. Forrest, 2012. [14] ................................................. 9 Figure 1.7: General depictions of four influential MHV designs. Adapted by permission from Springer Nature Customer Service Centre GmbH: Springer Nature Anticoagulation for Cardiac Prosthetic Devices: Prosthetic Heart Valves, Left Ventricular Assist Devices, and Septal Closure Devices by Matthew T. Crim, Supriya Shore, Suegene K. Lee et al 2018 [17] ......................................................................................................... 10 Figure 2.1: A visualization of the relationship between GOA and EOA in a BMHV. Adapted by permission from Elsevier. [Effective Orifice Area during Exercise in Bileaflet Mechanical Valve Prostheses], P. B. Bertrand et al, 30, 404-413, 2017. [40] ................................................................................... 17 Figure 3.1: From left to right, rendered views of the Okanagan Valve in the closed and open position, respectively. ..................................................................... 22 xi Figure 3.2: An isometric view of a leaflet is shown in drawing form. ............................... 24 Figure 3.3: The raised dome is shown through a cross-sectional view of the housing as a drawing. ................................................................................................... 25 Figure 3.4: The side view of the leaflet is shown as a drawing, with reference points denoted on the figure. In the closed position, the leading edge contacts the other leaflet and the trailing edge contacts the housing. ........................ 26 Figure 3.5: An isometric view of the housing is shown in drawing form. ......................... 27 Figure 3.6: Cross-sectional rendered view of the OKV. The raised dome on the housing which constrains the maximum opening angle of the leaflet is circled and labeled as point A. ........................................................................ 28 Figure 3.7: Cross-sectional view of the OKV in the closed position with the three contact points circled in red. The contact points from left to right are: between both leaflets, at the hinge between the leaflet and housing, and the trailing edge of the leaflet and the housing. ............................................ 29 Figure 3.8: View of the Hinge Area on a Closed OKV ........................................................ 30 Figure 4.1: A 3D modelled aortic root used for the exterior dimensions of the flow simulation. ....................................................................................................... 36 Figure 4.2: 83.0% Open Position Flow Simulation ............................................................ 36 Figure 4.3: The defined control volume considered around the valve. (Left) A is the leaflet tip, AD is the inlet, BC is the outlet, y is the distance between AD and the section of interest, EF is the section of interest, and phi is the angular stroke of the leaflet. (Right) A is the leaflet tip, B is the trailing edge of the leaflet, 0 is the pivot, O is the angle between 0B and 0E, 0B is the outlet, a is the major diameter of the housing, b is the minor diameter of the housing. ................................................................................. 37 Figure 4.4: The defined control volume considered around one leaflet in an arbitrary position. The control volume is considered from the outside. Moments and forces acting on the leaflet are shown by the red arrows. ...................... 39 Figure 4.5: An OKV prototype comprised of early revisions of both the saddle housing and curved leaflets fixed in place with pinned hinges. ..................... 42 Figure 4.6: An OKV housing printed out of ABS and smoothed using an acetone vapour process. The deformation of the housing and surface aberrations that occurred from briefly softening the material are depicted. Closeup View (Left), Full Housing (Right) ...................................................................... 43 xii Figure 4.7: A nearly finalized FFF OKV prototype. This prototype was printed by hand to give a representation of the final design's appearance. ............................ 43 Figure 4.8: The high-precision polymer Okanagan Valve prototype fixed in a custom gasket, atop a spare gasket. ............................................................................ 44 Figure 4.9: The aluminum Okanagan Valve prototype. The surface finish is characteristic of the DMLS process used to print this prototype. The pinned hinges that comprise the pivot mechanism of this prototype can be seen clearly in the left-hand image. ........................................................... 45 Figure 4.10: A dimensioned drawing of the gasket. All units are given in mm, the inner diameter is not shown as it is customizable for the valve being tested. ............................................................................................................. 46 Figure 4.11: TPU gasket prototype printed on a FFF 3D printer. The area of interest circled on the left-hand side demonstrates the layer separation that plagued gaskets created out of this material. ................................................ 47 Figure 4.12: Cross-sectional view of gasket mold V1. ......................................................... 48 Figure 4.13: Cross-sectional view of gasket mold V2. ......................................................... 49 Figure 4.14: Cross-sectional view of gasket mold V3. ......................................................... 50 Figure 4.15: The poor surface finish of the gaskets is the result of trapped air bubbles with the silicone as it set in the mold. This issue could be eliminated by degassing the silicon prior to filling the mold. ................................................ 50 Figure 4.16: The upper component closes flush against the top of the poured gasket, however, it does not impart any dimensions and can be removed to improve surface finish. .................................................................................... 51 Figure 4.17: Flow through a valve divided into the forward flow, closing volume, and leakage volume. The closing and leakage volume combined constitute the regurgitant flow through the valve each cycle [7].................................... 52 Figure 4.18: The Steady-Stage Leakage Analyzer at VSI. The high-precision polymer prototype OKV can be seen in the central orifice of the device fixed in position by its accompanying gasket. The gasket is identified by its pink colour. Not pictured is the upper component of the device which contains the fluid-filled silicone tube. ............................................................. 53 Figure 4.19: The ViVitro SuperPump apparatus. Left to right, the SuperPump pulse duplicator and its power and control system. Reproduced with permission from ViVitro Labs Inc. [73] ........................................................... 54 xiii Figure 4.20: The Pulse Duplicator at VSI Labs in Victoria, BC ............................................. 55 Figure 4.21: The Okanagan Valve Mounted in the Pulse Duplicator .................................. 56 Figure 4.22: The Pulse Duplicator with the OKV Affixed Within ......................................... 57 Figure 4.23: The Pulse Duplicator with Leonardo Attached Testing the OKV .................... 58 Figure 5.1: Tabulated results from the numerical analysis conducted on the OKV producing closing time, closing volume, maximum regurgitant blood flow velocity, and maximum leaflet tip velocity. Five different ovalities of the OKV and an SJM model as control were considered. The results are displayed in graphical form with the corresponding valve geometry. ........... 60 Figure 5.2: Hemodynamic performance results generated through in-vitro analysis of the aluminum OKV prototype in a pulse duplicator. Results represent five cycles of systole superimposed together. (A) Left-hand scale represents the valve open area, which is marked with five solid lines, one for each cycle. Right-hand scale represents the aortic valve flow rate which is marked by the dashed line. Apf denotes the peak forward aortic valve flow rate. A0f denotes when forward aortic valve flow rate reaches zero. (B) Left-hand scale represents the regional flow velocity, which is marked with five solid lines, one for each cycle. RBV represents regional backflow velocity; minimum, mean, and maximum values are depicted. ..... 63 Figure A.1: Upscaled OKV leaflet with multiple print defects. (Left) The exterior surface of the leaflet. (Right) The side view of the leaflet displaying the hinge geometry. .............................................................................................. 77 Figure A.2: The bottom view of an early full-scale OKV prototype, highlighting the pinned hinges. ................................................................................................. 77 Figure A.3: Two views of a full-scale OKV prototype; featuring improved print quality, functional hinges, and straight leaflets. ............................................. 78 Figure A.4: Two views of a full-scale OKV prototype; featuring a revised leaflet tip geometry, flat housing outlet, and a more aggressive inlet cutout. .............. 78 Figure A.5: Two views of a near-final revision full-scale OKV prototype. Printed using an FFF 3D printer, finished by hand, and painted to represent the appearance of a realistic BMHV. ..................................................................... 79 Figure A.6: Two views of a full-scale OKV housing prototype printed from PLA without a suture ring. ...................................................................................... 79 Figure A.7: TPU Gasket Layer Separation (Closeup View) ................................................. 80 xiv Figure A.8: OKV (left) and BMHV (right) Side Profile at 16.6% Open ................................ 81 Figure A.9: OKV (left) and BMHV (right) Side Profile at 33.2% Open ................................ 81 Figure A.10: OKV (left) and BMHV (right) Side Profile at 49.8% Open ............................... 82 Figure A.11: OKV (left) and BMHV (right) Side Profile at 66.4% Open ............................... 82 Figure A.12: OKV (left) and BMHV (right) Side Profile at 83.0% Open ............................... 82 Figure A.13: OKV (left) and BMHV (right) Side Profile at 100.0% Open ............................. 83 Figure A.14: OKV (left) and BMHV (right) at 16.6% Open ................................................... 83 Figure A.15: OKV (left) and BMHV (right) at 33.2% Open ................................................... 83 Figure A.16: OKV (left) and BMHV (right) at 49.8% Open ................................................... 84 Figure A.17: OKV (left) and BMHV (right) at 66.4% Open ................................................... 84 Figure A.18: OKV (left) and BMHV (right) at 83.0% Open ................................................... 84 Figure A.19: OKV (left) and BMHV (right) at 100.0% Open ................................................. 85 Figure A.20: Exterior Render of the Simulated Aorta (Front View) .................................... 86 Figure A.21: Exterior Render of the Simulated Aorta (Side View) ...................................... 86 Figure A.22: Four rendered side views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. .................................................. 87 Figure A.23: Four rendered front views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. .................................................. 88 Figure A.24: Four rendered top views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. .................................................. 89 Figure A.25: Four rendered bottom views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. ............................................. 90 Figure A.26: Four rendered bottom views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. ............................................. 91 xv List of Abbreviations ABS Acrylonitrile Butadiene Styrene ARF Acute Rheumatic Fever AV Atrioventricular BC British Columbia BPM Beats Per Minute BMHV Bileaflet Mechanical Heart Valve BHV Bioprosthetic Heart Valve CFD Computational Fluid Dynamics CL CryoLife DMLS Direct Metal Laser Sintering EOA Effective Orifice Area FDM Fused Deposition Modeling FFF Fused Filament Fabrication GOA Geometric Orifice Area HR Heart Rate HHS Human Heart Simulator HV Heart Valve LFS Low-Force Stereolithography MHV Mechanical Heart Valves MSA Mechanical Surface Area OKV Okanagan Valve PHV Prosthetic Heart Valve PLA Polylactic Acid PPM Patient Prosthesis Mismatch PTFE Polytetrafluoroethylene PVE Prosthetic Valve Endocarditis RBC Red Blood Cells RBV Regional Backflow Velocity RHD Rheumatic Heart Disease SAVR Surgical Aortic Valve Replacement xvi SJM St. Jude Medical SL Semilunar TAVR Transcatheter Aortic Valve Replacement THV Transcatheter Heart Valve TPU Thermoplastic Polyurethane UBC University of British Columbia VHD Valvular Heart Disease VSI ViVitro Systems Inc. xvii Acknowledgments I would like to take a moment to express my most sincere thanks to my supervisor Dr. Hadi Mohammadi for offering me the opportunity to explore my passion for Biomedical Engineering. I will always be grateful for the knowledge, support, and guidance he has offered me throughout my endeavor in graduate studies. Additionally, I would like to thank Lawrence Scotten from ViVitro Systems Inc. for sharing his knowledge of in-vitro heart simulators with me and for opening his lab and home to an inquisitive engineering student. Last, but not least, I would like to extend special thanks to my parents, sisters, friends, and lab mates for their endless love and support for my antics. I could not have finished my research without them steadfast by my side, yelling at me to keep writing. xviii Dedication To my mother and father, thank you for always supporting me, and for shaping me into the man I am today.Introduction and Background 1 Introduction and Background In this chapter, a general overview will be given of cardiac anatomy, valvular heart disease, and the historical development of prosthetic heart valves. The objective of this thesis is to develop a novel mechanical heart valve (MHV) that has comparable hemodynamics to those of a native heart valve. This would be a major improvement over current MHV designs as it would allow for safe implantation without requiring any post-operative medications. Such a MHV would be a true replacement for a failing native valve. As their designs mimic the native heart valves, current bioprosthetic heart valves (BHV) can be implanted without the need for anticoagulants unlike MHVs. However, while the average MHV will remain functional for a lifetime, a BHV will degrade over time and only has a usable lifespan up to approximately 20 years. As BHVs are typically manufactured from decellularized animal tissue, they do not self-repair as living tissue does and will continually degrade until failure. BHV lifespan is a serious consideration because the procedure for implantation, surgical aortic valve replacement, is a complex open-heart surgery that requires stopping a patientâs heart. Since such a procedure cannot be safely performed on elderly or very ill individuals, it is possible for a patient to outlive their prosthesis without the option of reoperation. Similar problems arise from anticoagulant reliance as they have considerable side effects such as how they can: react to other medications, complicate medical procedures, and increase the risk of internal bleeding. Patients and physicians are currently forced to weigh the pros and cons of each style of heart valve prosthesis with neither affording an ideal solution. A heart valve replacement that can outlast a patient without requiring any medication to support it is sorely needed to improve health care today. With this lofty goal in mind, this thesis and the design of the Okanagan Valve were initiated. Introduction and Background 2 1.1 Native Heart Valves The heart is arguably the most critical organ in the human body as it is responsible for pumping blood containing nutrients and waste products throughout the body. It is a two-stage pulsatile pump and will beat over two billion times during the average human lifespan. The heart is divided into left and right sides; with the right side being responsible for pulmonary circulation and the left side responsible for systemic circulation. The left side is larger and stronger than the right side due to the increased pressures and resistance it must overcome for systemic circulation. This correlates with the higher prevalence of left side valve conditions which will be further discussed in the following section [1]. Each side is comprised of an atrium, ventricle, semilunar (SL) valve, and atrioventricular (AV) valve. The primary function of the SL and AV valves is to maintain unidirectional flow through both halves of the heart and to prevent any regurgitation from occurring. Actuation of both AV and SL valves occurs solely due to the pressure gradient across the valve, and not from a muscular contraction. The location of the heart, the four heart valves, and the path of deoxygenated and oxygenated blood flow through the right and left halves of the heart can be seen in Figure 1.1. Figure 1.1: (A) Heart location in the human body (B) Labelled cross-sectional view of the human heart. Reproduced with permission from National Heart, Lung, and Blood Institute; National Institutes of Health; U.S. Department of Health and Human Services. [2] Introduction and Background 3 The two SL valves are named the pulmonary and aortic valve based on their positions in the right and left sides of the heart, respectively. The valves are located at the exits of their respective ventricles and their closure signifies the end of systole. Both SL valves share the same overall shape with three similarly sized cusps that close together to seal the valve. The cusps of the aortic valve differ from those of the pulmonary valve as two of them are just below the ostia of the two coronary arteries. These are known as the left coronary cusp and the right coronary cusp and are slightly larger than the third cusp, referred to as the non-coronary cusp. The shape of the aortic valve in both open and closed positions, as well as the coronary ostia (in the upper two-thirds), can be seen in Figure 1.2 as pictured looking towards the heart from the aorta. Figure 1.2: A healthy aortic valve in the open and closed positions. Modified with permission from Jill Rhead under a Creative Commons CC BY-NC-ND 4.0 license. [3] The two AV valves within the heart are known as the tricuspid valve and the mitral valve (also referred to as the bicuspid valve) and have a very different shape than the SL valves. The AV valves are situated between the atria and the ventricles on their respective sides of the heart. The tricuspid valve is on the right side and is comprised of three irregularly shaped cusps, whereas, the mitral valve is located on the left side and has two irregularly shaped cusps. The cusps of the AV valves are also connected to the interior of the ventricles by the subvalvular apparatus. The term subvalvular apparatus refers to fibrous structures known as chordae tendineae and papillary muscles that function to anchor the AV valve cusps to the ventricles, Introduction and Background 4 preventing prolapse into the atria. It is important to note that the subvalvular apparatus does not actuate the valve leaflets in any way, it only prevents prolapse from the higher pressures associated with ventricular contraction. Closure of the AV valves signifies the end of diastole. A cross-section of the left ventricle showing the mitral valve and the subvalvular apparatus is depicted in Figure 1.3, which also shows how the leaflets are anchored to the ventricle. Figure 1.3: The anatomy and major components of the mitral valve. Reproduced with permission from [4], Copyright Massachusetts Medical Society. 1.2 Valvular Heart Disease Valvular heart disease (VHD) is an umbrella term that covers a range of cardiovascular conditions that afflict human heart valves. Between 10-20% of surgical cardiac procedures in the United States are to address VHD [5]. The prevalence of VHD is estimated to be approximately 2.5% within the developed world [1]. Three-quarters of all cases of VHD in the developed world are caused by either mitral regurgitation or aortic stenosis [1]. This is dissimilar to the developing world where the majority of VHD is a result of bacterial infections Introduction and Background 5 (typically streptococcus or staphylococcus) that progress to acute rheumatic fever (ARF) [6]. The long-term effects of ARF on the heart valves, which can progress in severity with successive infections, are known as rheumatic heart disease (RHD) [6]. Although RHD progresses differently than the degenerative etiologies of VHD in the developed world, the cardiovascular damage sustained is predominantly to the aortic and mitral valves. As mentioned earlier, the valves on the left side of the heart are far more susceptible to damage because of the higher pressures they experience as a part of the systemic circulation. In an average individual, the pulmonary and tricuspid valves face pressures of approximately 30 mm Hg [7]. The aortic and mitral valves, however, must withstand pressures of approximately 100 mm Hg and 150 mm Hg respectively [7]. This additional loading accelerates the degradation of heart valves on the left side, whether they are native or implanted. 1.2.1 Stenosis and Incompetence The decreased function of a heart valve afflicted by VHD can be categorized as either stenosis, incompetence, or both. Stenosis is classified by damage to valve leaflets that results in a reduced lumen in the valve and obstructs blood flow through it. A stenotic valve will have a higher pressure drop across it and the leaflets will not actuate in a smooth fashion. Valvular incompetence occurs when damaged leaflets are unable to close fully resulting in large amounts of regurgitation. Stenosis and incompetence are typically caused by progressive thickening of valve leaflets and calcification, which occur naturally due to aging and is expected in individuals over the age of 65 [1]. Congenital defects such as bicuspid aortic valves can also result in stenosis and incompetence as well as accelerate the calcification process. In Figure 1.4 below, a healthy aortic valve is depicted alongside one suffering from severe calcification with clearly visible calcium nodules. Introduction and Background 6 Figure 1.4: (A) Healthy Aortic Valve (B) Severely Calcified Aortic Valve. Adapted by permission from Springer Nature Customer Service Centre GmbH: Springer Nature, Nature Reviews Disease Primers, [Calcific aortic stenosis], B. R. Lindman et al, 2016. [8] Unfortunately, there is no known pharmacotherapy that has been proven to be effective in reversing or limiting the progression of valve calcification [8]. This leaves surgical repair or valve replacement as the only viable options for patients suffering from stenosis or incompetence requiring intervention. For patients requiring heart valve replacements, there is a variety of different options that can be implanted with their individual pros and cons. 1.3 Non-Mechanical Heart Valve Protheses Non-mechanical heart valve prostheses are artificial heart valves made from decellularized animal tissue, either bovine or porcine in origin. Pericardial tissue is used to construct these valves and it is taken from the pericardium of the donor animal, which is a sac that surrounds the heart. To prevent rejection once implanted, the pericardial tissue is treated with glutaraldehyde ensuring biocompatibility. The irregularity of natural tissue also necessitates detailed scanning of donor pericardial tissue to ensure valves are constructed from material of uniform thickness. These valves are trileaflet in design, similar to the SL valves, and have excellent hemodynamics, comparable to those of the native valves they replace. One major benefit of the superior hemodynamic performance of these valves is that they can be implanted without any need for supplementary anticoagulant therapy. The major limitation of non-mechanical heart valve prostheses is the inevitable degradation of the tissue they are constructed from. They experience wear in the same manner as native valves, which is to say that the leaflets are thickened by fibrous growth and their movement restricted by calcification. Unlike native tissue which naturally repairs itself, these valves do Introduction and Background 7 not and are more susceptible to wear and tear under cyclic loading. As a result of this, over 50% of patients develop complications after 10 years and the majority must be replaced 20 years post-implantation [9]. The significance of this is that the surgical risk of heart valve replacement is even higher for patients undergoing reoperation or heart valve prosthesis replacement. Non-mechanical heart valve prostheses can be classified into two categories, bioprosthetic heart valves and transcatheter heart valves. The defining features of both types will be elaborated in the following two sub-sections. 1.3.1 Bioprosthetic Heart Valves Bioprosthetic heart valves (BHV) are valves made from animal pericardial tissue that are surgically implanted within patients. Surgical aortic valve replacement (SAVR) is the procedure performed to replace compromised aortic valves with BHVs and has a high surgical risk as it is an open-heart surgery. This requires cutting open the patientâs chest to expose the heart and placing the patient on a heart-lung bypass machine while the heart is stopped and the procedure is performed. BHVs come in two different forms, stented and stentless. Stented valves have leaflets fixed within a rigid stent which can then be implanted into most patients. Stentless valves are different such that the aortic root is replaced by a donor root with a valve within it. The major benefit of stentless BHVs is their large effective orifice area and the related low incidence rate of patient-prosthesis mismatch [10], [11]. The stented Medtronic Hancock II valve and the stentless Edwards Prima Plus valve are both depicted in Figure 1.5 below for comparison. Introduction and Background 8 Figure 1.5: (A) Stented Medtronic Hancock II BHV (B) Stentless Edwards Prima Plus BHV. Adapted by permission from BMJ Publishing Group Limited. [Transcatheter heart valve implantation for failing surgical bioprostheses: technical considerations and evidence for valve-in-valve procedures], D. Mylotte et al, 99, 960-967, 2013. [12] 1.3.2 Transcatheter Heart Valves Transcatheter heart valves (THV) are specially modified bioprosthetic valves that are implanted percutaneously through a catheter. THVs are a relatively new technology that is in the process of widespread adoption due to the benefits of transcatheter aortic valve replacement (TAVR) over SAVR. As TAVR is performed percutaneously, it is not necessary to stop a patientâs heart or to open the chest cavity to expose the heart. This significantly reduces the risks involved as well as the recovery time for the patient. As the patient population requiring heart valve replacement is typically elderly and often cannot undergo open-heart surgery, THVs is in some cases the only option available. Another benefit of THVs is they allow for the valve-in-valve (VIV) procedure to be performed. VIV is a procedure where a THV is used to replace a failing BHV that was implanted surgically and is otherwise inoperable due to the high associated mortality [13]. There are currently two forms of THVs that can be implanted, some that require a balloon catheter to expand a stainless-steel valve into position and others that utilize self-expanding nitinol frames. An example of the first variety is the Edwards-Sapien Valve. This valve requires a balloon catheter on which the valve is tightly crimped. Once in position, balloon valvuloplasty is performed, rapidly followed by balloon expansion of the valve into its final position [14]. The Medtronic CoreValve is an example of the second variety which utilizes the shape memory alloy nitinol. Once the valve is in position, it undergoes what is known as a Introduction and Background 9 martensitic transformation from the warmer body temperature and gradually expands into its final form [14]. Depictions of both THVs can be seen in Figure 1.6 below. Figure 1.6: (A) Edwards-Sapien Valve (B) Medtronic CoreValve. Reproduced with permission under a Creative Commons CC BY-NC-ND 4.0 license. "Transcatheter aortic valve replacement: design, clinical application, and future challenges." by J. K. Forrest, 2012. [14] 1.4 Mechanical Heart Valves Mechanical heart valves (MHV) are valves made entirely from biocompatible nonbiological material that are surgically implanted into patients. Due to their construction from rigid materials, these valves do not mimic native heart valves as BHVs do and their hemodynamic performance suffers accordingly. Typically constructed from pyrolytic carbon, these valves do not suffer from material failure as BHVs do and are a permanent solution [7]. However, their poor hemodynamic performance forces reliance on anticoagulant therapy to counteract their thrombogenicity [7], [9], [15]. MHVs today are fundamentally incompatible with the human body because of their harsh effects on blood. The valves induce large shear stresses on red blood cells (RBC) causing platelet activation to occur. Platelet activation is a serious concern as it leads to the formation of thrombi which can progress to an embolism, stroke, or heart attack. While medication can be prescribed to successfully counteract this shortcoming, heavy reliance on anticoagulants can have adverse effects on patients, aggravate other pre-existing conditions, and interact with foods and medications [16]. Introduction and Background 10 MHVs have taken on many design architectures since their concept was first introduced in 1952 by Charles Hufnagel. Four of the most influential design styles can be seen in Figure 1.7 below, and the defining points of each will be elaborated in the following sub-sections. Figure 1.7: General depictions of four influential MHV designs. Adapted by permission from Springer Nature Customer Service Centre GmbH: Springer Nature Anticoagulation for Cardiac Prosthetic Devices: Prosthetic Heart Valves, Left Ventricular Assist Devices, and Septal Closure Devices by Matthew T. Crim, Supriya Shore, Suegene K. Lee et al 2018 [17] 1.4.1 Caged Ball MHVs The first MHV to be implanted in a patient was Hufnagelâs Valve, a caged ball design, invented by Dr. Charles Hufnagel in 1952 [9]. This initial design had many shortcomings, such as its implantation location in the descending aorta, but was a successful proof-of-concept for MHVs. The most widely used valves featuring a caged ball design were the Edwards-Starr models, which were implanted over 300,000 times over the course of their production. While their performance left much to be desired, caged ball valves proved to be a reliable solution for many years. Introduction and Background 11 1.4.2 Caged Disc MHVs A slight variation on the caged ball design was the caged disc design, also known as single leaflet valves. These were designed to reduce the amount of regurgitation and improve the response time by having the disc close flush with the housing thanks to the large surface area of the disc. Unfortunately, the flat surface of the disc was a significant impedance to blood flow through the valve resulting in a high rate of thromboembolisms [9]. 1.4.3 Tilting Disc MHVs Tilting disc MHVs incorporated clever engineering of a retaining cage such that the motion of the disc was controlled through the opening and closing phases. This accurate control provided a significant improvement in hemodynamics over previous designs. Unfortunately, the discs continued to occlude the lumen through the valve, increasing the pressure drop across it. Later revisions of the tilting disc design rotated the disc further during the opening phase so that in the open position the disc would minimally block the lumen through the valve [7]. Tilting disc designs greatly improved upon the caged ball design and they were used throughout the world from 1960 through 1977 when the St. Jude Medical Bileaflet valve was first introduced [9]. 1.4.4 Bileaflet MHVs The most commonly implanted MHVs worldwide today are variations of the bileaflet mechanical heart valve (BMHV) design. The two most popular BMHVs in the world today are CryoLife (CL) On-X valve and the gold standard St. Jude Medical (SJM) Regent valve. These designs are based on a common architecture of a housing acting as the annulus with two pivoting leaflets housed within [7]. The first design to use this architecture was the Kalke-Lillehei bileaflet valve in 1964 which failed because of the lack of an appropriate biomaterial. The St. Jude Medical bileaflet valve was first introduced in 1977 and was the first MHV to utilize a fully pyrolytic carbon construction. This pure carbon structure allowed for the valve to be implanted with minimal biocompatibility issues. Its bileaflet design permitted the leaflets to open to over 80 degrees, significantly improving the lumen and effective orifice area of the valve in comparison to previous designs. One of the highlights of its design was its unique, hourglass-shaped recessed hinge which effectively controlled the opening and Introduction and Background 12 closing angles of the leaflets which protruded into the recess. Since 1977, variations of the St. Jude Medical valve have been implanted in millions of patients worldwide and they are the most successful MHVs to date [9]. Nearly all MHVs implanted today are either the SJM Regent valve or the CL On-X bileaflet valve. Despite their excellent reliability, both designs still require the constant use of anticoagulants to combat their poor hemodynamic performance. In the next chapter, an in-depth overview of the limitations of BMHVs will be given as a performance baseline for the next generation of BMHV design. 1.5 Thesis Outline This thesis details the limitations of modern BMHVs, the design of the Okanagan Valve (OKV), and the methods and results from evaluating the OKV as a potential new design architecture for BMHVs. Including this introductory chapter on valvular heart disease and prosthetic heart valves, this thesis is comprised of six chapters. The outline of the following chapters is as follows: â¢ Chapter 2 (Design Limitations of Bileaflet Mechanical Heart Valves): Describes the limitations of modern bileaflet mechanical heart valves and aspects that are believed to be critical for creating a replacement design architecture. â¢ Chapter 3 (Okanagan Valve): Describes the design of the OKV through its main components: the leaflets, the housing, and the hinges. Critical design elements of the OKV are defined in detail. â¢ Chapter 4 (Methods): Describes the OKV prototypes that were manufactured and the four methods that were used to evaluate the performance of those prototypes. â¢ Chapter 5 (Results): Describes the results of the methods used to evaluate the performance of the prototypes. â¢ Chapter 6 (Conclusions and Future Works): Summarizes the main conclusions of this thesis, the work completed, the limitations of the analysis, and potential avenues for advancing this research.Design Limitations of Bileaflet Mechanical Heart Valves 13 Design Limitations of Bileaflet Mechanical Heart Valves In this chapter, the design limitations of BMHVs and the critical parameters for creating a successful BMHV will be discussed. In its basic function, an MHV is a check valve ensuring unidirectional flow through it. Complications arise in designing a heart valve to not have adverse reactions with blood, the immune system, or suffer from premature failure. The following sections will elaborate on optimizing hemodynamic performance, durability, and biocompatibility of the prosthesis. 2.1 Hemodynamics In the context of biomedical engineering and this thesis, hemodynamics refers to âthe physical study of flowing blood and of all the solid structures (such as arteries) through which it flowsâ [18]. It is important to understand hemodynamics in order to control the thrombogenicity of MHVs. As discussed in the preceding chapter, anticoagulants are required to offset the poor hemodynamic performance of current MHVs following implantation. Without medication such as vitamin K antagonists, MHVs would rapidly cause the formation of thrombi which could manifest as a thromboembolism, stroke, or heart attack [16]. Thrombus formation induced by a prosthetic heart valve (PHV) can be explained by the components of Virchowâs triad: surface factors, hemodynamic factors, and hemostatic factors [19]â[21]. Hemostatic factors have a reduced role in PHV thrombosis as they are typically patient specific factors for hypercoagulability, however their effects must be considered when evaluating high-risk patients for an appropriate prosthetic [19]. Hemodynamic and surface factors serve a larger role in PHV associated thrombi formation, both of which will be discussed in the following subsections. The hemodynamic performance of a heart valve can be characterized by four primary criteria: shear stresses, stagnant flow, regurgitation, and effective orifice area. Design Limitations of Bileaflet Mechanical Heart Valves 14 2.1.1 Shear Stresses High mechanical shear stress patterns directly adjacent to and within an MHV are believed to be a major contributing factor to RBC damage [22]. These elevated shear stresses can rapidly cause lethal damage to RBCs (hemolysis) leading to thrombus or pannus formation. There are multiple known causes for the violent shear stresses seen in modern MHVs, namely leaflet location, rapid closure, and hinge gap width [23]â[25]. This can be characterized by evaluating the Reynolds shear stress (RSS) of an MHV. The thresholds for hemolysis and platelet activation are conservatively estimated at 1500 dynes/cm2 and 100 dynes/cm2 respectively [26]. The location of leaflets in current MHVs directly contributes to their poor hemodynamics. The leaflets are semi-circular in shape and are hinged centrally in a BMHV. This places them in the center of the valveâs lumen during the open phase. The leaflets are then obstructing flow at the center of the opening, directly where blood flow velocities are the greatest. This placement also leads to flow stagnation and separation, which are covered in the following subsection. To mitigate this behavior, a leaflet should not occlude the central region of the valve during the open phase. As blood flow out of the ventricles slows down at the end of systole, native heart valves begin to close softly. MHVs however, do not begin to close until the very end of systole and close very rapidly [15]. This can lead to elevated regional backflow velocities and the formation of âvelocity jetsâ that have been identified to have thrombogenic potential [15]. The ideal leaflet behavior would be for closure to begin as the end of systole approaches, and for the leaflets to be fully closed when systemic and ventricular pressures equalize as seen in native heart valves. Hinge gap width in the context of BMHVs refers to the distance between the ears of the leaflet and their corresponding hinge recess within the valve housing. In a 2014 study by Yoganathan et al., three variations of the St. Jude Medical BMHV were assessed for their thrombogenic potential [23]. The study found that the hinge gap width had a significant influence on the shear stress characteristics of the valves during the closed phase, under leakage flow. It was found that the largest hinge gap width resulted in the highest shear stress magnitudes [23]. Design Limitations of Bileaflet Mechanical Heart Valves 15 It was also determined that the smallest hinge gap width demonstrated the poorest hinge washout flow. This is a critical factor as flow stagnation within the hinge can lead to the accumulation of activated platelets and begin the coagulation cascade. The study concluded that the hinge gap width needs to be optimized for any given BMHV to reach a balance between shear stress and stagnant flow. 2.1.2 Flow Stagnation High concentrations of thrombogenic material, such as activated platelets, can accumulate in regions of stagnant flow. This non-physiological flow, which can be found in some regions of a BMHV, can lead to the further activation of platelets and thrombosis [7], [21]. Vortices can also occur at the outlet of BMHVs as the limited EOA flowing into the larger diameter of the aortic root can result in an adverse pressure gradient [27]. As blood components can accumulate and become trapped in these vortices, they can induce platelet activation and accumulation as with flow stagnation [28]. One common location where stagnant flow can be found in BMHVs is the hinge area [9], [23]. Thrombus formation is occasionally seen on the hinges of BMHVs and if left untreated, can progress along the housing towards its orifice, restricting the movement of the leaflet. The butterfly-shaped hinges characteristic of BMHVs place the female portion on the housing of the prosthetic, where there is a limited volume of blood flow. BMHV designs must demonstrate adequate wash of their hinge areas to ensure that the thrombogenic risk posed by flow stagnation is addressed. Manufacturers of BMHVs have taken to adjusting their hinge designs to improve the wash resulting in different designs, such as the smooth leaflet ear in the St. Jude Medical (SJM) valve and the sharp ear of the CarboMedics valve [29]. Studies have shown the improved hemodynamics of some hinge designs over others, indicating the existence of an optimal hinge design [29]. 2.1.3 Regurgitation Regurgitation in a BMHV is the result of retrograde flow through the valve as it closes as well as leakage through the valve while it is shut. These two components are known as the closing volume and the static leakage, respectively [7], [9]. As an excessive volume of regurgitant Design Limitations of Bileaflet Mechanical Heart Valves 16 flow will necessitate an increase in heart rate to maintain the cardiac output of the heart, total regurgitation should ideally match the performance of a native aortic valve. BMHVs are known for their large closing volumes, for example, the St. Jude Medical Regent valve (25 mm) has a closing volume of 11.2 mL/ beat [7]. The static leakage is approximately 450 ± 14 mL/min [23]. BMHVs are not an improvement on their predecessors in this regard, as both tilting disc and caged-ball MHVs have smaller regurgitant volumes on average [7], [9]. It is hypothesized that the large closing volume is related to the non-physiological closing action of BMHVs. While native and bioprosthetic heart valves begin to close as forward flow through the valve decelerates, BMHVs do not begin to close until retrograde flow begins to accelerate [30]. This behavior compounds with the 33 ms closing time and 55-60 degree travel arc of the SJM Regent leaflets [31], [32]. The static leakage in a BMHV is a result of the inter-leaflet gaps and the gap between the leaflets and the housing. Additionally, the allowance between the moving components of the hinge allows for static leakage retrograde through the valve. The gaps between the components of the valve are a necessity for ensuring reliable, frictionless actuation of the valve and cannot be eliminated entirely. The allowance within the hinge structure, however, is optimized such that adequate retrograde wash of the hinge area can be achieved during the closing phase [23], [29]. This wash also results in the formation of leakage flow velocity jets. Velocity jets are another important aspect of regurgitant flow in BMHVs, as they exhibit high shear stresses and have been observed in previous studies [25], [33]â[37]. Leveraging the computational and numerical methods developed in those previous studies would aid the development and refinement of a novel hinge balancing adequate wash, induced shear stresses, and leakage volume. 2.1.4 Effective Orifice Area It is critical for a heart valve prosthetic to have a large effective orifice area (EOA) relative to its geometric orifice area (GOA) so that it may be an ideal match for patients, mitigating patient prosthesis mismatch (PPM). PPM is a phenomenon where the largest PHV that can be safely implanted, based on the size of the patientâs aorta and the prosthetics GOA, is Design Limitations of Bileaflet Mechanical Heart Valves 17 insufficient for their needs. When the EOA of an implanted prosthetic is too small relative to a patientâs body size the result is poor hemodynamics, lower survival rates, and reduced left ventricular hypertrophy regression [38]. A meta-analysis of PPM following aortic valve replacement found survival rates were significantly improved in patients with no PPM when compared to those with moderate or severe PPM [39]. While the GOA of a PHV is fixed, the EOA is smaller and dependant on the design of the valve and its fluid interaction. Figure 2.1 below depicts both the EOA and GOA of a BMHV, demonstrating how the centrally hinged leaflets impede flow through the valve reducing its EOA. Figure 2.1: A visualization of the relationship between GOA and EOA in a BMHV. Adapted by permission from Elsevier. [Effective Orifice Area during Exercise in Bileaflet Mechanical Valve Prostheses], P. B. Bertrand et al, 30, 404-413, 2017. [40] Mitigating this limitation of BMHVs currently requires an aortic root enlargement procedure to accept a larger, better matched prosthetic. While this can often be performed safely, it does introduce risks to the procedure [41]. In some cases, these risks outweigh the dangers associated with PPM and the inadequate prosthetic is accepted as the best possible solution [38], [39]. Design Limitations of Bileaflet Mechanical Heart Valves 18 2.1.5 Cavitation While the opening mechanics of a BMHV are quite similar to those of a native or BHV, the closing mechanics are not. BMHVs begin closing as the retrograde flow accelerates, whereas native valves begin closing earlier as the forward flow decelerates [30]. The result is that BMHVs have a shorter closing time, approximately twice as fast, but a greater leaflet velocity than their native counterparts [30]. A consequence of this greater leaflet velocity is cavitation [42], [43]. Cavitation occurs when vapor bubbles form on the trailing surface of the leaflets as they close, and implode once the leaflets close completely, damaging the leaflet. This phenomenon occurs due to a local pressure drop on the leaflet surface below the vapor pressure of the working fluid [44]. The second major contributing factor to the cavitation observed with MHVs is known as squeeze flow [30], [42]â[45]. Squeeze flow, in the context of a BMHV, refers to the volume of working fluid that is rapidly displaced from the gap between the leaflet tip and the housing as it approaches the closed position [45]. Together, squeeze flow and leaflet velocity can result in the formation of cavitation bubbles in MHVs, and the effects are positively correlated with elevated heart rates [44]. There are two major adverse effects of cavitation in MHVs; their thrombogeneic potential, and accumulating damage caused to the prosthetic itself. The thrombogenecity of cavitation bubbles is due to both the high-velocity jets they induce and the high local pressures. The velocity jets from squeeze flow have been recorded up to a maximum velocity of 30 m/s, nearly an order of magnitude larger than the closing velocity [43], [45]. The maximum velocities are short bursts but the high shear stresses induced have a clear thrombogeneic potential. The prothrombotic potential of cavitation bubbles has been previously documented by several studies both in-vitro and in-vivo [46]â[49]. Cavitation is also damaging to the structure of MHVs. The implosion of cavitation bubbles can erode the leaflets by causing pitting [44]. This pitting can negatively affect the hemodynamics of a MHV by eroding coatings, roughing the polished finish, and by the thrombotic potential of the pits themselves. The pitting also accumulates over time, gradually increasing in severity [44]. Considering that BMHVs are designed and marketed for their durability and reliability, erosion should be Design Limitations of Bileaflet Mechanical Heart Valves 19 eliminated as it could result in degrading performance of the prosthetic long-term and premature failure. 2.2 Durability The defining feature of MHVs is their nearly indefinite lifespan which proves to be their primary advantage over BHVs. Most BMHVs are manufactured using pyrolytic carbon, either coated onto a graphite substrate or machined from a monolithic slab [9], [50]. Unlike the decellularized tissue that BHVs are made from, carbon does not degrade within the human body under normal circumstances. However, a very small portion of implanted BMHVs have failed prematurely due to structural failures [51]. The most common failure is leaflet escape, where a leaflet will separate from the prosthesis and migrate along the vascular system [51], [52]. As there is an extremely low occurrence rate of these failures, there is essentially no concerns associated with the durability of BMHVs [53]â[55]. Some forms of pyrolytic carbon, like boron alloyed pyrolytic carbon used by TRI-Technologies, have displayed higher failure rates than competitors, such as Carbomedicsâ Pyrolite® [50], [52]. Zhang et al. theorized that BMHVs manufactured from a graphite-pyrolytic carbon composite are more resistant to crack propagation than homogeneous pyrolytic carbon [50]. Due to the brittleness of the material, novel BMHV designs should be rigorously tested in-vitro to diagnose material failure due to stress concentrations. The current design of BMHVs places excessive stress on the hinges as the maximum opening angle of the leaflets is solely constrained by hinges. Potentially, there is room for improvement if those stresses can be distributed across the leaflet and the housing together, further increasing the durability of the prosthetic. Due to the high durability of current valves, the added benefit of such a modification would be difficult to ascertain. 2.3 Material Biocompatibility The biocompatibility of a BMHV is a very important factor in assessing its long-term performance as a PHV. As blood contacts biomaterials, reactive mechanisms activate which can negatively affect the performance of prosthetics and propagate into serious clinical events [56]. Modern MHVs have, for the most part, addressed the material biocompatibility Design Limitations of Bileaflet Mechanical Heart Valves 20 issues their predecessors suffered from regarding both the structure of the prosthetic, and the suture ring used for attachment. These improvements were enabled by advances in material science and selection for the construction of MHVs and their suture rings, reducing their thrombogenicity and improving their performance. 2.3.1 Valve Structure Since their first introduction, a wide variety of materials have been used for the construction of MHVs [9]. Many polymers were evaluated to determine their biocompatibility, but the majority of them suffered from thrombosis within hours [57]. A material was required that was rigid enough to construct an MHV, but also resistant to thrombus formation. Following experiments by Gott et al. that uncovered a method to bond heparin, an anticoagulant, to a substrate, pyrolytic carbon was rediscovered [57]. Initially developed by Jack Bokros as a coating for nuclear fuel rods, a sample of pyrolytic carbon was sent to Gott et al. to evaluate its performance [57], [58]. While attempts at bonding heparin to its surface were unsuccessful, pyrolytic carbon was found to be the least thrombogenic non-heparinized rigid material and remains critical for the fabrication of BMHVs today [9], [57], [58]. Pyrolytic carbon is the primary material used in nearly all BMHVs [9], [58]. It is usually applied as a coating on top of a graphite substrate through a vapor deposition process [59]. In its current state of development, it is unquestionably the superior material for constructing an MHV. Some studies have shown that its hemodynamic performance and thromboresistance can be improved by the application of diamond or superhydrophobic coatings [26], [60], [61]. However, the durability and stability of these coatings needs to be further investigated before they can be integrated into MHV designs. 2.3.2 Suture Ring Since their first introduction, the material used to create suture rings for BMHVs has changed. Originally fabricated from a trademarked form of polyester named Dacron, they were eventually changed to be manufactured from polytetrafluoroethylene (PTFE) to improve their biocompatibility. The PTFE suture rings were found to improve on the paravalvular leakage that occurred with Dacron suture rings [62]. PTFE, commonly known as Teflon, comprises the suture rings of the most widely used BMHVs today. Design Limitations of Bileaflet Mechanical Heart Valves 21 The primary area of research concerning suture rings today is combatting prosthetic valve endocarditis (PVE), which is an infection that can occur following heart valve replacement. While it is not a common infection, mortality rates associated with PVE have been reported as high as 75% [63]. One example of a suture ring developed to address PVE is the Silzone St. Jude Medical valve [64]. The unique silver coated suture ring was incorporated for its antibacterial protection. However, preliminary results reported a high incidence rate of both paravalvular leaks and thromboembolic events, and SJM elected to voluntarily recall the valve [63]â[67]. Other attempts have also been made to produce an antibacterial suture ring through coatings and saturation, however, there has been no definite solution to the problem [68]â[70].Design of the Okanagan Valve 22 Design of the Okanagan Valve In this chapter, the overall design of the Okanagan Valve will be detailed along with key design aspects of its major components: the leaflets, housing, and hinges. The following subsections will focus on the critical design elements that were regarded throughout the design process and the effects they are intended to achieve. The OKV is a departure from the design architecture that can be seen across BMHVs implanted today, as well as in the gold standard of MHVs, the St. Jude Medical BMHV. This is immediately apparent when viewing the OKV from any angle and can be seen in the rendered images in Figure 3.1 below. Figure 3.1: From left to right, rendered views of the Okanagan Valve in the closed and open position, respectively. The OKV design is comprised of two leaflets fixed within a housing with a suture ring attached to the exterior of the housing. The shapes of the leaflets and the housing are vastly different from other existing BMHVs and define the OKV. Their unique design enables the leaflets to open away from the center of the valve with the maximum opening angle constrained by multiple contact points. Similarly, the closed position of the leaflets is also constrained at multiple points. The multiple points of contact are intended to mitigate stresses induced on Design of the Okanagan Valve 23 the hinges during actuation of the valve. There are no additional components in the OKV design compared to other BMHVs. For the purpose of this thesis and testing, the OKV was designed around a 25mm inner diameter. The other components were dimensioned around the fixed inner diameter and comparable values found in commercial BMHVs. This inner diameter value was chosen arbitrarily as it is approximately the midpoint size for commercial BMHVs. It is anticipated that further revisions of the OKV would be evaluated at a full range of sizes to confirm their performance individually. The housing body is 1.35mm thick for the majority of its shape, however, its thickness varies in certain areas. The thickest points on the housing are the hinge locations. The parallel, flattened hinge surfaces are spaced 23mm apart across their width. These areas on the housing are 2.35mm thick at their thickest point. The thinnest points on the housing are where the leaflets close against the housing, these areas on both sides come to an edge which is anticipated to be modified following extensive in-vitro analysis. The suture ring modelled around the OKV is a semicircular ring with a 1.2mm radius that follows the exterior surface of the saddle-shaped housing. Overall, the housing is 30mm across at its widest and 10.2mm high at its tallest points. The leaflet dimensions consist of a 1mm thickness for much of their body but this reduces to a thickness of 0.65mm at the hinge areas. The widest point on the leaflets is across the hinges, measuring at 22.3mm wide. The height of the leaflets is set at 16mm. 3.1 Leaflet Design The curved leaflets of the OKV enable them to open away from the center of the housingâs orifice and nearly flush with the inner diameter of the housing. The curve of the leaflets is designed to maximize the orifice area and to minimize disturbance to the flow through the valve during systole. An isometric view of a leaflet can be seen in Figure 3.2 below. Design of the Okanagan Valve 24 Figure 3.2: An isometric view of a leaflet is shown in drawing form. The leading edge of the leaflet, the contact point between the two leaflets, is rounded to minimize the contact patch, resulting in a linear contact patch between the two leaflets. This was incorporated into the design to reduce the probability of red blood cells (RBC) being crushed between the two leaflets during actuation. For the same reason, the maximum opening angle of the leaflets is also constrained by small hemispherical features on the housing. These raised points create a single contact point between the housing and the leaflets, contacting them at their midpoint towards the leading edge. This point of contact between the leaflets and the housing can be seen in Figure 3.3 below. These features will be further described in the following subsection regarding the housing. Design of the Okanagan Valve 25 Figure 3.3: The raised dome is shown through a cross-sectional view of the housing as a drawing. As visualized in Figure 3.2 above, the female portions of the butterfly hinges are located on the leaflets. This is a departure from all modern BMHVs which place the male portion of the butterfly hinge on the leaflet. The reasoning for this change is to place the hinge opening closer to the center of the valveâs orifice. It is theorized that a more thorough wash of the hinges can be achieved by leveraging the greater blood velocities found nearer to the center of the valve. The faster velocities at the center are a result of the developing flow through the valve and the boundary layer resistance on the inner surface of the valve. Figure 3.4 below is a close-up view of the female hinge portion on the leaflet. Design of the Okanagan Valve 26 Figure 3.4: The side view of the leaflet is shown as a drawing, with reference points denoted on the figure. In the closed position, the leading edge contacts the other leaflet and the trailing edge contacts the housing. 3.2 Housing Design The housing of the OKV was designed to feature a saddle-shaped profile to accommodate the unique actuation of its leaflets. The saddle-shape of the housing also reflects the native aortic annulus more accurately than the cylindrical housings of traditional BMHVs. An isometric drawing of the housing can be seen in Figure 3.5 below. Design of the Okanagan Valve 27 Figure 3.5: An isometric view of the housing is shown in drawing form. This geometry can also be leveraged to reduce the surface area of the prosthesis to aid the OKVâs biocompatibility by reducing the amount of foreign material being implanted. Because of the housingâs (and its accompanying suture ringâs) similarity to the native aortic annulus, surgical implantation of the OKV is anticipated to be a less complex procedure than current BMHVs. A rendered image of the OKV with an area of interest circled can be seen in Figure 3.6 below. Design of the Okanagan Valve 28 Figure 3.6: Cross-sectional rendered view of the OKV. The raised dome on the housing which constrains the maximum opening angle of the leaflet is circled and labeled as point A. At point A in the image above, a hemispherical dome can be seen. This is the supplementary point of contact for the leaflets when they are in the open position. These domes prevent the leaflets from opening flush with the interior of the housing which would result in an undesirable planar contact surface. A planar contact surface would be a high-risk area for RBC lysis, hemolysis, to incur because of the large surface area in which RBCs could be crushed. The stresses induced by constraining the maximum opening angle of the leaflets are shared between the hinges and the domes on the housing. This was incorporated into the design in order to mitigate the risks of fatigue failure and to improve the lifespan of the OKV. The maximum closing angle of the leaflets is constrained in three different locations to also mitigate the risks of material fatigue failure. These three points are: between the leading edges of the leaflets, between the leaflet trailing edge and the housing, and the hinges themselves. In Figure 3.7 below, a cross-sectional view is shown with these three contact locations circled in red. Design of the Okanagan Valve 29 Figure 3.7: Cross-sectional view of the OKV in the closed position with the three contact points circled in red. The contact points from left to right are: between both leaflets, at the hinge between the leaflet and housing, and the trailing edge of the leaflet and the housing. It is theorized that these additional contact points for the maximum and minimum opening positions will effectively eliminate any concerns of fatigue failure by extending the lifespan of the prosthetic beyond the average human lifespan. 3.3 Hinge Design For its leaflet hinges, the OKV utilizes a similar shape to the âbutterflyâ hinges that have been a defining feature of BMHVs for decades. The hinges on the OKV retain the same âbutterflyâ geometry but relocate the female portion of the hinge to the leaflet and the male portion to the housing. This change to the hinge was made possible considering the geometry of the OKV leaflets. A female socket would be prohibitively difficult to incorporate onto the flat leaflets of standard BMHVs. The reasoning behind this modification to the hinges is to eliminate the need for a cavity in the housing as it is difficult to achieve adequate wash there. By moving the socket portion closer to the center of the valveâs lumen, where blood flow Design of the Okanagan Valve 30 velocities are greater, it is theorized that a more thorough wash of the hinge can be achieved. This is desirable to reduce the possibility of a stagnant region of blood near the hinges which can result in thrombosis. Figure 3.8 below depicts a drawing view of the hinge assembly as it appears on the valve in the closed position. Figure 3.8: View of the Hinge Area on a Closed OKV 3.4 Critical Design Elements The end goal for designing the OKV is to match the hemodynamics of a native valve while either meeting or exceeding the durability of the current gold standard of BMHVs, the St. Jude Medical Bileaflet Heart Valve. The following subsections detail specific attributes that were the driving forces for the design of the OKV. These elements were implemented because they were perceived to be beneficial in improving the performance of the valve. 3.4.1 Central Blood Flow Considering that the largest flow velocities through a pipe are found towards its center, it was desired to develop the OKV geometry such that blood flowed unobstructed through a single, central orifice. This was achieved by the novel design of the leaflets which causes them to Design of the Okanagan Valve 31 open towards the valve housing, away from the center of the annulus. This single, central flow orifice reduces disturbance to the blood flow through the valve in comparison to the three orifices found in modern BMHVs. Specifically, the disturbance typically induced by leaflets positioned in the path of blood during systole is eliminated. The OKVâs unique leaflet design results in its single orifice and contributes to other critical design elements such as the effective orifice area. 3.4.2 Effective Orifice Area As discussed in section 2.1.4, it is important for a PHV to have a large EOA relative to its GOA so that it may provide adequate hemodynamics for the patient, improving their long-term outcome [39]. The single central orifice of the OKV is believed to provide an improvement in EOA over traditional BMHVs as the leaflets no longer obstruct flow through the center of the valve. Free from the obstruction of the leaflets, visualized in Figure 2.1, the OKV should have a larger EOA-GOA ratio than any other BMHV. This may correspond with improved patient survival rates by limiting PPM, reducing the need for aortic root enlargement, and enabling implantation in biologically smaller patients. 3.4.3 Hinge Washing To prevent thromboembolic complications, it is important for mechanical heart valves to have their hinge areas washed out thoroughly during each cycle. The âbutterflyâ hinge design that has become commonplace in current designs is partially comprised of cavities in the housing that are prone to flow stagnation. To counteract this, the allowance between the âearâ protrusion from the leaflet and the housing cavity is carefully controlled, and both the housing and the cavity are designed to draw blood through them with each cycle of the valve. A significant correlation has been demonstrated between the geometric aspects of a butterfly hinge and its thromboembolic potential by previous studies [23], [35]. The function of clearing stagnant RBCs from the hinge is known as the wash or washout of a MHV. For traditional BMHVs, drawing flow through the hinge area alongside the housing of the valve for an effective wash is difficult due to the limited volumetric flow through that area. To address this in the design of the OKV, the female portion of the hinge was moved to the leaflet with the male portion extending from the housing and the hinge itself was enlarged. Design of the Okanagan Valve 32 Both of these modifications are intended to improve wash in the OKV on their own as well as catalyzing the effect of the other. The unique geometries of the OKV allow for the hinge area to be larger than in traditional BMHVs where the leaflet âearâ is limited in thickness to the thickness of the leaflet itself. The anticipated result of enlarging the hinge area is that it will increase the effect any features designed to improve wash have on the flow through the valve, as well as increasing the volume of blood washing through the hinge. Flipping the hinge orientation brings the hinge recess to an area with a relatively greater volumetric flow of blood and the possibility of drawing blood through the interior face of the recess on the leaflet, across the housing protrusion in the hinge, and out alongside the housing interior. This flow pattern is just one possibility of how blood flow could interact with the hinge for the OKV based on the design of the surrounding area. An exact design for the hinge area was deemed beyond the scope of this thesis, and the need for further study is elaborated as future works in Chapter 6: Conclusions and Future Works. 3.4.4 Surface Area The biocompatibility of an implanted prosthesis is a function of both the material and the total surface area implanted within the patient. In the case of MHVs, pyrolytic carbon has been the primary material used for their fabrication since the 1970âs due to its excellent material properties and biocompatibility [58]. As there is no replacement material for pyrolytic carbon at the time of this thesisâ writing, it is anticipated that it will be utilized for the fabrication of the OKV. In order to improve the biocompatibility of the OKV, the total surface area of the prosthetic must be minimized to reduce the amount of foreign material that is implanted within the patient. Reducing the total implanted surface area should translate to reduced immunorejection of the prosthetic, therefore improving the prognosis of any prospective patients. The saddle shaped housing of the OKV features a smaller surface area than traditional cylindrical housings. The curved leaflets however, have an increased surface area in comparison to their flat counterparts. Design of the Okanagan Valve 33 Measurements were taken of a 25 mm ID model of the OKV reproduced in SolidWorks to compare the mechanical surface area (MSA) to those of the SJM and CL BMHVs. MSA values for the SJM and CL BMHVs were taken from a study conducted by Cho et al. [71]. For the 25 mm sizes of the SJM and CL BMHVs, the MSAs were measured to be 17.64 cm2 and 22.28 cm2, respectively. The MSA of the latest version of the OKV is measured to be 27.32 cm2 which is 54.88% larger than that of the SJM valve and 22.62% larger than that of the CL valve. Due to the constantly evolving nature of the design, the total surface area is not yet finalized and will be improved. It should be noted that despite the increased surface area of the leaflets, the reduction in the housing surface area is anticipated to result in a net reduction in total implanted MSA. 3.4.5 Shear Stress The primary complication following implantation of a BMHV is the occurrence of life-threatening thromboembolic events. These events are primarily believed to be the result of high mechanical shear stresses found within the implanted prosthetics [22]. In order to counteract the high shear stresses BMHVs induce, patients are prescribed anticoagulants which they must take for the remainder of their lives. The major contributing design aspect of BMHVs that results in their high induced shear stresses is not agreed upon. The current consensus is that it is a function of both the leaflets obstructing the core flow and the jet velocities that occur due to the geometry of the valves during opening and closure. These factors consistently result in shear stresses found both experimentally and analytically that exceed the threshold for platelet activation and starting the coagulation cascade [72]. To reduce the reliance on anticoagulants, the OKV was designed to address both the leaflet obstruction and jet velocity induced shear stresses. The leaflets of the OKV were designed to open outwards in the direction of the housing, creating a single central orifice for blood to flow through. It is understood from fluid dynamics that as flow develops in a pipe, the fluid velocity will become greatest near the center and theoretically zero near the walls. It is inferred that this change to the leafletsâ opening mechanism will significantly reduce the shear stresses that are induced during the forward flow phase relative to traditional BMHVs. Design of the Okanagan Valve 34 The degree to which these shear stresses are reduced will require an extensive in-vitro analysis of the OKV using a pulse duplicator. The high shear stresses induced by jet velocities are a result of the small orifice openings seen in BMHVs just following leaflet opening and just prior to leaflet closure. The pressure difference which causes valve actuation and the small open orifices combine to produce temporary spikes in fluid velocity which are known to cause platelet activation [72]. As traditional BMHVs open, a relationship can be drawn between each degree of leaflet movement and the three open orifices created through the valve; one central between the leaflets and two flanking the leaflets. If the same relationship is drawn for the OKV, it is apparent that the single large opening will increase in size faster than the three openings of a traditional BMHV for each degree of leaflet motion. Since platelet activation from high shear stresses is also a function of the exposure time, it can be inferred that the single orifice of the OKV will mitigate the issue of jet velocities and the resultant shear stress spikes. 3.4.6 Closing Volume During the closing phase of a MHV, a small volume of blood is regurgitated before the leaflets can fully close. This is known as the closing volume of the valve. Due to the passive nature of MHVs, this closing volume is difficult to eliminate as the leaflets only begin to close after the outlet (aortic) pressure exceeds the inlet (ventricular) pressure. In order to reduce the closing volume of the OKV, its range of motion was minimized by increasing the length of the leaflets. This change resulted in a smaller range of motion compared to traditional BMHVs and shortened both the opening and closing phases of the valve. The reduction in the closing phaseâs duration is anticipated to translate directly to a reduction in the closing volume of the OKV and can be evaluated both computationally and experimentally. Results 35 Methods In this chapter, the methods used to simulate and evaluate the performance of the Okanagan Valve will be discussed. The four methods used were flow simulation, numerical modeling, steady-state leakage testing, and in-vitro testing. The latter two methods differ from the former two as they were physically conducted on multiple OKV prototypes at ViVitro Systems Inc. (VSI) in Victoria, BC using FDA approved in-vitro equipment. The flow simulation tool was used to provide a quick visual representation of flow patterns around the OKVâs leaflets. This provided a proof of concept that the valve geometry was behaving as intended prior to additional testing. Next, a numerical analysis was conducted to estimate the performance of the OKV by studying one leaflet during a cardiac cycle. To experimentally corroborate those numerical results, in-vitro testing equipment was utilized to test the OKV in both steady-state and pulsatile configurations. The following subsections describe the aforementioned methods in greater detail. 4.1 Flow Simulation A quick flow simulation was performed using SolidWorks Flow Simulation to illustrate the flow patterns through the OKVâs novel leaflet shape. A second flow simulation was performed using semi-circular leaflets akin those found in modern BMHVs to be used as a control. These simulations were performed as part of a term project for MECH 533 â Biofluids at the Vancouver campus of UBC. The project was structured to produce results that would benefit the studentâs thesis. As SolidWorks Flow Simulation is incapable of dynamically modeling the leaflets in motion, six different simulations were conducted for each valve structure, equally spaced between the nearly closed and fully open positions. The valve structure consisted of a 25mm cylinder as an inlet and an approximated aortic root as an outlet. In Figure 4.1 below, the housing used for the simulation is depicted. Results 36 Figure 4.1: A 3D modelled aortic root used for the exterior dimensions of the flow simulation. Realistic hinges were not used in this simulation for the sake of simplicity. Instead, the leaflets were mated using simple pivots and constrained to the correct opening and closing angles. In Figure 4.2 below, a 3D view of each valve simulation can be seen at the 83.0% open position demonstrating flow through the valves. Figure 4.2: 83.0% Open Position Flow Simulation The variables used in these simulations were greatly simplified as the main purpose was to produce graphics that visually demonstrated the single, central flow orifice of the OKV in comparison to industry-standard BMHVs. Due to its inherent simplicity, the simulations do not provide any confirmation or validation of the performance characteristics of the OKV and will not be examined later in Chapter 5: Results. For further review, all images generated from the flow simulations will be included in the appendices under Appendix B: Flow Simulation. Results 37 4.2 Numerical Modeling In order to simulate the regurgitant flow from the aortic root into the left ventricle, a numerical model was applied to create an in-silico study of the behavior. The theorized hemodynamic improvements of the Okanagan Valve are characterized by the regurgitation volume and velocities of both blood through the valve and the leaflet tip during the closing phase. The finite strips method was used for the numerical model to solve the equations of motion using the 4th order Runge-Kutta method. The numerical model was compiled as custom code written in Fortran. The regurgitation flow into the left ventricle was simulated by applying a half-symmetrical model of the valve (one leaflet). It is assumed that the ventricular (Pv) and aortic pressures (Pao) are equal and constant prior to the start of the closing phase [9]. A control volume is defined around the valve model and can be seen in Figure 4.3 below. Figure 4.3: The defined control volume considered around the valve. (Left) A is the leaflet tip, AD is the inlet, BC is the outlet, y is the distance between AD and the section of interest, EF is the section of interest, and phi is the angular stroke of the leaflet. (Right) A is the leaflet tip, B is the trailing edge of the leaflet, 0 is the pivot, O is the angle between 0B and 0E, 0B is the outlet, a is the major diameter of the housing, b is the minor diameter of the housing. OBA (effectively a 3D curve) shows the leaflet pivoted at point O. Points AD and BC refer to the inlet and outlet, respectively, and the section of interest is EF. The blood velocities at EF, AD, and BC are VEF, VAD, and VBC respectively. AD and EF are separated by distance âyâ. It is Results 38 assumed that the gradient of velocities and pressures from the inlet to the outlet remain linear. The blood velocity of the arbitrary section, VEF, is calculated with respect to the inlet and outlet blood velocities, VAD and VBC respectively [9]. This is accomplished by applying the continuity equation to the control volume with moving boundaries as follows: a AD (ðð´ð· â ðð¡ð´ð·) = a EF (ðð¸ð¹ â ðð¡ð¸ð¹) +ððððð¡ (1) ðð¡ð´ð· = ðð´ ð ððð ð (2) ðð¡ð¸ð¹ = ðð¸ ð ððð ð (3) ðð =(ð´ð· + ð¸ð¹ ) 2 a ð¸ð· ð ððð (4) VtAD, VtBC, and VtEF denote the velocity of the leaflet tips at the inlet, outlet, and arbitrary section respectively. The major diameter of the elliptic housing is âaâ, the angular motion of the leaflet is ð, and the velocity of the control volume is ðð. The reaction force acting on the leaflet, Freac, is calculated by solving Equations (1) â (4). Freac is also calculated using an alternative method as follows. The control volume is considered the same as in the previous method and there are two forces acting on both the inlet and the outlet as shown in Figure 4.4. The forces acting on the inlet AD are AD ððð and ?Ì?ð´ð·ðð´ð·; the forces acting on the outlet BC are BC ðð£ and ?Ì?ðµð¶ððµð¶. Results 39 Figure 4.4: The defined control volume considered around one leaflet in an arbitrary position. The control volume is considered from the outside. Moments and forces acting on the leaflet are shown by the red arrows. Figure 4.4 shows that the four forces acting on the inlet and outlet must be in equilibrium with Freac as well as the momentum change in the control volume, ððððð¡. This is shown as: ðð¡ðµð¶ = âððµ ð cos ð (5) ?Ì?ðµð¶ = ð a BC (ððµð¶ â ðð¡ðµð¶) (6) ?Ì?ð´ð· = ð a AD (ðð´ð· â ðð¡ð´ð·) (7) (a AD ððð â a BC ðð£ â ð¹ðððð) â (?Ì?ðµð¶ððµð¶ â ?Ì?ð´ð·ðð´ð·) =ððððð¡ (8) The mass of the control volume and its velocity, M and V respectively, are calculated as: ð = ð a ð¶ð· ð´ð· + ðµð¶2sin ð (9) ð =ðð´ð· + ððµð¶2 (10) In Equations (6), (7), and (9), ï² denotes the blood density. Freac, previously calculated by solving Equations (1) â (4), is now calculated by solving Equations (5) â (10). The following Results 40 energy equations for a control volume with moving boundaries are solved to calculate the final velocity and pressure values as shown [9]: ðð´ð·2 â ðð¸ð¹22= â«ðððð¡ðð¦ âðð´ð· â ðð¸ð¹ðð (11) ðððð¡= [(ðð)ð¡ â (ðð)ð¡âð¿ð¡]/ð¿ð¡ (12) â«ðððð¡ð0 ðð¦ = â[(ðð)ð¡ â (ðð)ð¡âð¿ð¡]ð¿ð¡ ð¿(ð§â) sin ð ðð=1 (13) ð¿(ð§â) =ðð (14) This modeling strategy utilized the pressure domain including the aortic and ventricular pressures to determine the inlet and outlet velocities. This is contrary to conventional models in which the waveform and velocity profile are directly attributed to the inlet. Using Equation (15), which expresses the dynamic motion of the leaflet, the angular position of the leaflet with respect to time is calculated. The pressure-induced moment is Mp, the gravity-induced moment is Mg, and the angular mass of the leaflet with respect to the hinges is Jp. The angular acceleration is depicted as ?Ì?. The angular position of the leaflet is calculated as follows: ðð + ðð = ð½ð?Ì? (15) ððð = ð´ðððð ð¿ð¡ (16) ðð = â â« (ðð¸ð¹ âðððð´ð·) a ðð¸ ðð = â(ðð¸ð¹ â ðð´ð·) a ðð¸ ð¿ððð=1 (17) ðð = ðð ( ð/2) cos ð (18) The mass of a single leaflet is denoted as m. The leafletâs motion begins based on an initial velocity of 0 m/s from its maximum opening angle which is set to 89º. The regurgitation flow is assessed based on Equation (16), in which Aor denotes the total orifice area. Results 41 4.2.1 Modelling Setup The objective of this numerical modeling setup is to estimate the regurgitation flow volume of a complete cardiac cycle. The time increment, δt, is set to 0.05 ms. As Freac is calculated through two different approaches, it is assumed that convergence occurs when the deterrence between the numerical values for Freac becomes less than or equal to 0.01 Newtons. The number of strips, N, was chosen to be 45 and the Runge-Kutta method of the 4th order was implemented in order to solve the differential equations. Heart rate (HR) was set to 70 beats per minute (bpm) and the cardiac output was set to 6.1 L/min. The average aortic pressure was set to nominal systolic pressure of 120 mmHg. 4.3 Okanagan Valve Prototypes Throughout the design process of the Okanagan Valve, numerous prototypes were made of varying quality based on the available rapid prototyping equipment. Early in the design process, crude models were printed using a basic fused filament fabrication (FFF) 3D printer to help visualize the valve in its current form and aid the iterative design process. FFF is sometimes referred to as the trademarked term fused deposition modeling (FDM), however, the more accurate term, FFF, will be used for the contents of this thesis. The 3D printer initially used for prototype fabrication was a Monoprice Select Mini V2 printer. The primary printing material used was polylactic acid (PLA). To work around the limited print resolution of the printer, many initial prototypes were printed at an enlarged scale. The limited complexity of early prototypes also reduced the difficulty of achieving successful prints. Figure 4.5 below is an example of a very early prototype that was used to better understand the mechanics of the valve. Results 42 Figure 4.5: An OKV prototype comprised of early revisions of both the saddle housing and curved leaflets fixed in place with pinned hinges. The FFF 3D printing process results in layer lines across the surface of a finished print due to how the thermoplastic material is extruded and cooled for the deposition of each layer. These layer lines can be seen along the sides of the prototype in Figure 4.5 above. The rough quality is also visible along the top edge of the prototype. In an attempt to improve the surface finish, acetone was utilized through multiple approaches with varying success. Acetone has a slight reaction with PLA making it slightly malleable, however, it has a drastic smoothing effect on acrylonitrile butadiene styrene (ABS). Submerging prints in acetone had an immediate effect on prints, however, it was too great of an effect for the desired goal. Brushing prints with acetone had a similar effect but merely substituted brush strokes for the layer lines. The only method that was slightly successful was placing ABS prototypes in a sealed container with acetone. The acetone vapors gradually smoothed the print surfaces to a glossy finish. However, this consistently resulted in warped prints which did not function correctly due to the strict tolerances involved. As FFF was proving to be too inaccurate for HV prototype fabrication, further attempts at acetone smoothing were abandoned. An example of a smooth but warped OKV housing can be seen in Figure 4.6 below. Results 43 Figure 4.6: An OKV housing printed out of ABS and smoothed using an acetone vapour process. The deformation of the housing and surface aberrations that occurred from briefly softening the material are depicted. Closeup View (Left), Full Housing (Right) As the design of the OKV became more complex, FFF prototypes were printed for use as visual aids with no expectation of functionality. This necessitated a heavier reliance on support structures, careful calibration of the printer itself, and a higher incidence rate of print failure. Despite these challenges, prototypes of the OKV were continually produced using the FFF printer until late in the design process. Figure 4.7 below depicts a nearly finalized model of the OKV that was printed using the FFF printer and painted by hand. Figure 4.7: A nearly finalized FFF OKV prototype. This prototype was printed by hand to give a representation of the final design's appearance. Results 44 Once the valve design had matured to the point where it was ready for in-vitro testing, a high-precision polymer prototype was 3D printed in the STAR Labs located at UBC Okanagan. This polymer prototype was printed on a Connex500 3D printer and can be seen held within a custom silicone gasket in Figure 4.8 below. These custom gaskets were designed and manufactured to dimensions provided by VSI to affix HV prototypes within the in-vitro testing equipment. Their design procedure will be further elaborated in the following subsection. Figure 4.8: The high-precision polymer Okanagan Valve prototype fixed in a custom gasket, atop a spare gasket. In the image above, the characteristic lines of layer-based polymer 3D printing can be clearly seen on the back of the leaflet. Two metal pins were used for each leaflet to fix them to the housing while allowing a single degree of freedom so that they could pivot freely. Despite the best efforts of those involved in manufacturing this prototype, the material unfortunately lacked adequate attributes for in-vitro testing. The poor durability and stiffness of the material led to its structural failure shortly into the testing procedure. This prototypeâs performance will be discussed further in Chapter 5: Results. Results 45 A replacement for the polymer prototype was created following its failure and was also manufactured using a 3D printing process. This prototype was printed out of aluminum (ALSi10Mg) by a third-party fabricator outside of UBC, Additive Metal Manufacturing. It was printed on an EOS M290 printer using the direct metal laser sintering (DMLS) process. Due to resolution limitations associated with DMLS printing, some minor modifications were made to the valve design to ensure successful fabrication. Overall, the function and design of the prototype were preserved. To create the hinge mechanism for the aluminum prototype, 3D printed holes on the housing and leaflets were drilled and honed before steel pins were press-fit inside to fix them together. The contact surfaces of the hinges, as well as some aberrations from the print process, were filed smooth to improve function. In Figure 4.9 below, the aluminum prototype is shown from two different angles and the textured surface finish from the DMLS process can be seen clearly. Figure 4.9: The aluminum Okanagan Valve prototype. The surface finish is characteristic of the DMLS process used to print this prototype. The pinned hinges that comprise the pivot mechanism of this prototype can be seen clearly in the left-hand image. The aluminum prototype improved on the shortcomings of the polymer prototype but still exhibited some flex when placed in the heart simulator. This model also lacked the butterfly hinges which are critical to the design of an MHV, due to the technical limitations of the manufacturing process. The implications of the hinge change will be further discussed in the limitations section of the final chapter. The performance of the aluminum prototype will be further discussed in the following chapter. Results 46 4.4 In-Vitro Mounting Gaskets In order to utilize the in-vitro testing equipment, gaskets for securing the OKV prototypes needed to be manufactured. These gaskets would have to externally match the mounting plates on the VSI hardware and internally hold the OKV prototype in place. Due to the unique shape of the OKV, specifically its saddle-shaped housing, existing gaskets at VSI could not be used or retrofitted. Initially, it was theorized that the gaskets could be manufactured from flexible 3D printed materials. The benefits of this approach would be the speed and ease with which gaskets could be created. Thermoplastic polyurethane (TPU) was selected as a possible material for evaluation as it was compatible with the available 3D printing hardware. The gasket was modelled in SolidWorks with the appropriate dimensions and can be seen in Figure 4.10 below. Figure 4.10: A dimensioned drawing of the gasket. All units are given in mm, the inner diameter is not shown as it is customizable for the valve being tested. After printing multiple models in TPU with modified parameters such as infill, vertical shell thickness, and horizontal shell thickness, it was determined that a gasket made from TPU would not be functional. While TPU elastically deforms more than other materials such as Results 47 ABS and PLA, it is not elastic enough to form a good seal and cannot deform enough to accept a HV prototype within. When under stress, the TPU prototypes have a tendency for their layers to separate from one another and for the exterior shell to collapse. One of the surviving TPU prototypes is shown in Figure 4.11 below. The gap visible on the left-hand side of the TPU prototype is an example of the aforementioned layer separation. Figure 4.11: TPU gasket prototype printed on a FFF 3D printer. The area of interest circled on the left-hand side demonstrates the layer separation that plagued gaskets created out of this material. At the suggestion of Lawrence Scotten from VSI, a silicone mold material with a 30A shore hardness was selected for further gasket prototypes. Smooth-On MoldMax 30A was chosen for these prototypes as its material specifications met the requirements of the in-vitro hardware. In order to produce gaskets from the silicone material, it was necessary to design molds. Initial molds suffered from unnecessarily complex designs and it was not until the fourth iteration that gaskets could be produced reliably. The initial mold design featured six different 3D printed components, from two different materials, as well as fastening hardware. The final iteration of the mold design reduced that to two printed components from one material and no fastening hardware. The initial mold design featured top, bottom, and core components printed out of PLA which would be press-fit together. Washers printed out of TPU would seal the core against the top and bottom, as well as the top and bottom together. Finally, the mold would be secured using Results 48 the mounting flanges on the top and bottom components and filled through a hole in the top. A render of this design is depicted in Figure 4.12 below. Figure 4.12: Cross-sectional view of gasket mold V1. Two observations were made upon assembling the mold; firstly, the fastening hardware was unnecessary as the press-fit components formed a tight seal, and secondly, the fill hole was too small. Upon removing cured silicone from this mold, it was noted that the TPU washers did not have the desired sealing effect and had instead created additional areas for surface defects on the final product. The next revision was to rectify these identified issues by removing the flanges, the reliefs for the TPU washers, and increasing the fill hole diameter. A render of this design is depicted in Figure 4.13 below. Results 49 Figure 4.13: Cross-sectional view of gasket mold V2. The gasket produced by this mold was an improvement on the last but had surface defects caused by the gap left between the core and both the top and bottom components. The next revision addressed the defects left by those gaps by incorporating the core into the bottom component of the mold. The top component was also split into two halves to aid removal of the cured silicone. A render of this design is depicted in Figure 4.14 below. Results 50 Figure 4.14: Cross-sectional view of gasket mold V3. This third iteration of the mold design produced functional gaskets, however, the lack of access to a vacuum chamber was evident in the surface finish. Small cavities caused by air bubbles pockmarked the top surface of the gasket. Occasionally, a large trapped air bubble would result in a large cavity along the top surface of the mold. An example of the surface finish left by these air bubbles can be seen in Figure 4.15 below. Figure 4.15: The poor surface finish of the gaskets is the result of trapped air bubbles with the silicone as it set in the mold. This issue could be eliminated by degassing the silicon prior to filling the mold. The final iteration of the mold design did not involve any additions to the design, but the removal of the top component. Figure 4.16 below is a cross-sectional view of the mold design which shows how the top component does not contribute to any geometrical features. By Results 51 removing the top from the mold apparatus, the internal cavity could be easily filled to the brim of the mold, which allowed for a more consistent gasket. Figure 4.16: The upper component closes flush against the top of the poured gasket, however, it does not impart any dimensions and can be removed to improve surface finish. 4.5 In-Vitro Steady-State Leakage Testing As discussed previously, MHVs are manufactured from rigid materials and their operation can be described as leaflets actuating within a housing to predetermined maximum opening and closing angles. For the aortic valve, leaflet actuation occurs due to the pressure difference between the left ventricle and the circulatory system through systole and diastole. The motion of the rigid leaflets within the housing necessitates allowances between their external dimensions to minimize frictional losses and the resulting pressure drop across the valve. It is important to minimize friction as the leaflets must never become stuck in any position whether it be open, closed, or in between. An excessive pressure drop across the valve (aortic) will result in left ventricular hypertrophy over time and a stuck leaflet can have fatal consequences for a patient. Complications arise from the necessary allowances as they are gaps between the leaflets and housing through which blood can âleakâ when a HV is closed. This is known as static leakage volume and is a component of a HVâs transvalvular regurgitation along with its regurgitant closing volume. The static leakage volume is visualized in Figure 4.17 below as the area marked âleakage volumeâ. Results 52 Figure 4.17: Flow through a valve divided into the forward flow, closing volume, and leakage volume. The closing and leakage volume combined constitute the regurgitant flow through the valve each cycle [7]. In order to test the leakage flow of an MHV, an apparatus is required that can hold the valve in place and exert systemic pressure on the exit side of the valve for a set period. The pressure exerted can be varied to test various physiological conditions. An example of such a device is the steady-state leakage analyzer at VSI and is pictured in Figure 4.18 below. Results 53 Figure 4.18: The Steady-Stage Leakage Analyzer at VSI. The high-precision polymer prototype OKV can be seen in the central orifice of the device fixed in position by its accompanying gasket. The gasket is identified by its pink colour. Not pictured is the upper component of the device which contains the fluid-filled silicone tube. The device pictured above is the apparatus that was used to test the leakage flow and basic operation of an early Okanagan Valve prototype. The device is comprised of a base with user controls and a reservoir, above which rests a fixture for holding the HV being tested. Not pictured is a third fluid-filled silicone tube that attaches to the top of the apparatus and extends upwards to create the necessary head. For our testing purposes, a pressure of 120/80 mmHg was applied. The working fluid was saline with a viscosity of 1 mPaâs and a density of 1.0 g/mL. To ensure correct fitment within the apparatus, any HV being tested must have an accompanying gasket made from silicone. This gasket has identical external dimensions to the gasket required for fixing an HV into the pulse duplicator. This gasket can be easily Results 54 identified in Figure 4.18 by its bright pink colour. Fixed within the gasket is the high-precision polymer Okanagan Valve prototype. 4.6 In-Vitro Pulse Duplicator Testing In-vitro testing of a HV prototype is conducted with a specialized piece of equipment known as a pulse duplicator. Often referred to as a human heart simulator, a pulse duplicator mimics the pulsatile flow of the heart by recreating the heartâs chambers and pumping water through them at variable pressures and volumetric flow rates. Pulsatile flow is generated by the pump component, named the SuperPump, that forces water through the Left Model Heartâs chambers. Together, the Left Model Heart and the SuperPump comprise the pulse duplicator system. Figure 4.19 below depicts the modern, digital ViVitro SuperPump and its power and control system which combined, produce the user-selected waveform for testing. Figure 4.19: The ViVitro SuperPump apparatus. Left to right, the SuperPump pulse duplicator and its power and control system. Reproduced with permission from ViVitro Labs Inc. [73] A wide range of valves including stented, stentless, percutaneous, and mechanical can be accommodated into the aortic and pulmonary positions by installing them with a bespoke silicone gasket. These gaskets are identical to the ones used in steady-state leakage testing, conforming to both the shape of the pulse duplicator housing and to the prosthetic within. The transparent construction of the Left Model Heart allows for clear viewing of the valve during testing and various ports around the device enable data collection. The pulse duplicator is powered by the ViVitro SuperPump and records data to a local computer on their proprietary software. Results 55 The pulse duplicator that was used for testing the Okanagan Valve in Victoria can be seen in Figure 4.20 below and is different from the current commercially available version. It is an older version of the pulse duplicator that uses analog control systems unlike the digital systems of the current pulse duplicators. This changes how the pulse duplicator is operated but does not affect the systemâs performance nor does it affect the validity of the results generated in any way. Figure 4.20: The Pulse Duplicator at VSI Labs in Victoria, BC The Okanagan Valve was mounted in the aortic position for testing and the pulse duplicator was operated by its co-inventor, Lawrence Scotten. A key component of the system, which Results 56 contains the simulated aortic root, is missing from Figure 4.20 above. It was removed to expose the mounting area for the MHV. In Figure 4.21 below, the aluminum Okanagan Valve prototype is visible mounted into the aortic position within the distinct pink gasket. The high-precision polymer prototype was also mounted in the pulse duplicator but suffered a mechanical failure shortly after testing began. The performance of the high-precision polymer prototype will be further discussed in Chapter 5: Results. Figure 4.21: The Okanagan Valve Mounted in the Pulse Duplicator The mounting area for the MHV features a slot that loosely fixes the gasket in position above the ventricular outlet as shown in Figure 4.21 above. The MHV is then securely fixed into its position by attaching the aortic root component to the pulse duplicator which also completes Results 57 the fluid loop of the system by providing a return line to the fluid reservoir. Figure 4.22 below shows the complete pulse duplicator system with the aortic root component affixed. Figure 4.22: The Pulse Duplicator with the OKV Affixed Within The final component of the pulse duplicator system is the Leonardo apparatus, developed by VSI Labs. Leonardo is comprised of a light source and a photosensor which attach above and below the HV being evaluated. By measuring the amount of light that passes through the HV during operation, Leonardo allows for the instantaneous volumetric flow rate through the Results 58 valve to be calculated at any given time. Combining this data with valve motion data enables further insight into the tested valveâs performance and the regional backflow velocities [74]. The Leonardo apparatus can be seen in Figure 4.23 below, attached to the pulse duplicator. Figure 4.23: The Pulse Duplicator with Leonardo Attached Testing the OKV The photo in Figure 4.23 above was captured during the in-vitro testing of the aluminum OKV prototype in Victoria, BC. The performance of the aluminum prototype will be further discussed in Chapter 5: Results.Results 59 Results In this chapter, the results of the methods used to evaluate the performance of the Okanagan Valve will be discussed. The results obtained from numerical modelling, steady-state leakage testing, and in-vitro testing will be discussed. As mentioned in Chapter 4: Methods, the flow simulation results are omitted as they do not accurately model the characteristics of the valve beyond basic validation. The results from each flow simulation have been included within Appendix B for further reading. The steady-state and in-vitro testing results were generated from testing conducted on multiple valve prototypes at VSI. 5.1 Numerical Results In the numerical analysis, the derived equations were leveraged to calculate values for four different parameters; blood flow velocity (regurgitant), leaflet tip velocity, closing volume, and closing time. The closing volume was calculated by considering the relative velocity between the leaflet tips and the nearby blood flow. Five different ovalities were considered for the OKV housing and leaflets in order to determine the optimal ovality. The results from this numerical analysis can be seen tabulated in Figure 5.1 below. Results 60 Figure 5.1: Tabulated results from the numerical analysis conducted on the OKV producing closing time, closing volume, maximum regurgitant blood flow velocity, and maximum leaflet tip velocity. Five different ovalities of the OKV and an SJM model as control were considered. The results are displayed in graphical form with the corresponding valve geometry. The five different ovalities utilized for numerical analysis are denoted as Oval 1 through Oval 5. Oval 1 featured a minor diameter of 24 mm and a major diameter of 26 mm. Oval 2 featured a minor diameter of 23 mm and a major diameter of 27 mm. Oval 3 featured a minor diameter of 22 mm and a major diameter of 28 mm. Oval 4 featured a minor diameter of 21 mm and a major diameter of 29 mm. Oval 5 featured a minor diameter of 20 mm and a major diameter of 30 mm. A 25 mm ID SJM BMHV model was also used as a control. The closing time for the OKV was calculated to be between 17-21 ms, dependent on the ovality used. The SJM control was calculated to have a 33 ms closing time. It should be noted that the closing time of the SJM valve is known to vary based on the size of the valve in question, and typically ranges from 26-33 ms. The maximum regurgitant blood velocity for the OKV was calculated to be between 1.7-1.9 m/s, dependent on the ovality used. The SJM control was calculated to have a maximum regurgitant blood velocity of 2.7 m/s. It should be noted that these maximum blood velocities do not account for the presence of velocity jets which can occur due to squeeze flow in the Results 61 SJM valve. The presence of squeeze flow and the resultant velocity jets with the OKV have not been determined at this point. A lower regurgitant blood velocity is indicative of a softer closing phase for an MHV. This suggests a reduction in squeeze flow, if present at all. The maximum leaflet tip velocity for the OKV was calculated to be between 1.6-1.8 m/s, dependent on the ovality used. The SJM control was calculated to have a maximum leaflet tip velocity of 2.3 m/s. Leaflet tip velocity is associated with the formation of low-pressure regions trailing the leaflet tip, which can induce cavitation if the pressure drops below the fluid vapor pressure. The closing volume for the OKV was calculated to be between 5.3-6.4 cc, dependent on the ovality used. The SJM control was calculated to have a closing volume of 10.5 cc. The closing volume was calculated by considering the relationship between regurgitant blood velocity and closing time. These results show that compared to all other considered ovalities, the Oval 2 model performed favorably. Further evaluation of additional ovalities with greater similarity to Oval 2 is warranted to determine the optimal ovality for an MHV. It is important to note that every ovality considered for the OKV produced favorable results compared to the control SJM model. 5.2 In-Vitro Steady-State Leakage Results In-vitro steady-state leakage testing was conducted on two different prototypes of the OKV. The high-precision polymer prototype and the aluminum prototype were both tested under the same conditions. The pressure applied in both tests was 120/80 mmHg, with a saline solution as the working fluid. The viscosity and density of the saline solution were 1 mPa·s and 1.0 g/mL, respectively. Testing of the prototypes was conducted to determine their geometric leakage area, prior to conducting pulsatile tests. A geometric leakage area range deemed ideal for pulsatile testing was between 0.02-0.04 cm2 [15]. The high-precision polymer prototype was found to have a geometric leakage area of 0.03 cm2. The aluminum prototype was found to have a geometric leakage area of 0.04 cm2. These Results 62 results fell within the desired range for the prototypes and they were cleared to proceed with pulsatile testing. It should be noted that both prototypes utilized inaccurate hinge mechanics. Due to the limitations of prototype manufacturing, pins were used for the hinges of both prototypes to facilitate actuation of the leaflets. This change to the pivoting mechanism required allowances between the leaflets and housing that are not representative of the intended final design of the OKV. This limitation will be further discussed in the following chapter under 6.2: Limitations. 5.3 In-Vitro Pulse Duplicator Results In-vitro pulse duplicator testing was conducted on two different prototypes of the OKV. Both the high-precision polymer prototype and the aluminum prototype were tested with the same conditions. The pulse duplicator was configured to reproduce average physiological conditions. These conditions are represented by heart rate, cardiac output, and pressure and were configured to 70 bpm, 5 L/min, and 120/80 mmHg. The working fluid was a saline solution with a viscosity and density of 1 mPa·s and 1.0 g/mL, respectively. In-vitro analysis of the high-precision polymer prototype was unsuccessful as it failed due to leaflet escape during the testing procedure. Prior to testing, the prototypeâs leaflets actuated freely and did not appear too fragile for the process. The prototype was fixed within a gasket and installed in the simulated aortic annulus without issue. The pulse duplicator was activated at a reduced heart rate and cardiac output than the desired conditions and steadily increased. Prior to reaching the desired conditions, the hinge area on one leaflet fractured and the leaflet separated from the valve. Data was not recorded as the prototype was never actuating under the desired physiological conditions. The leaflet was damaged beyond repair and spare leaflets had not been fabricated for a potential replacement. Following these results, the high-precision polymer was deemed too fragile for in-vitro analysis and a replacement was sought out. The aluminum OKV prototype was produced for in-vitro testing as it was believed to be capable of withstanding the forces incurred during the analysis procedure. In-vitro analysis Results 63 of the aluminum prototype was successful, and data was recorded during the experimental procedure. Five cycles of systole were recorded for analyzing the performance of the OKV and the data has been graphed below in Figure 5.2. Figure 5.2: Hemodynamic performance results generated through in-vitro analysis of the aluminum OKV prototype in a pulse duplicator. Results represent five cycles of systole superimposed together. (A) Left-hand scale represents the valve open area, which is marked with five solid lines, one for each cycle. Right-hand scale represents the aortic valve flow rate which is marked by the dashed line. Apf denotes the peak forward aortic valve flow rate. A0f denotes when forward aortic valve flow rate reaches zero. (B) Left-hand scale represents the regional flow velocity, which is marked with five solid lines, one for each cycle. RBV represents regional backflow velocity; minimum, mean, and maximum values are depicted. Results 64 In Figure 5.2 (A), the valve open area (cm2) and aortic valve flow rate (mL/s) are graphed with respect to time (ms). The valve open area for the five cycles of systole that were recorded are shown as solid lines superimposed upon one another. The aortic valve flow rate is shown as a dashed line. Apf and A0f represent the peak forward aortic valve flow rate and when forward aortic valve flow rate reaches zero, respectively. The closing volume can be determined by calculating the area under the curve when the aortic valve flow rate is negative. The experimental closing volume was determined to be approximately 5 cc. The experimental closing time for the valve is approximately 35 ms. The closing time places the OKV prototype between those of BMHVs and bioprosthetic valves which are faster and slower, respectively. The closing volume for bioprosthetic heart valves is on average <2 cc and 11.2 cc for a 25 mm SJM Regent bileaflet valve [7]. By comparison, the experimentally recorded closing volume of approximately 5 cc for the OKV shows a clear improvement over the current gold standard. In Figure 5.2 (B), the regional flow velocity (m/s) is plotted with respect to time (ms). Data from the five recorded systolic cycles is shown superimposed together. The minimum, mean, and maximum values of the regional backflow velocity (RBV) were determined to be 41 m/s, 44 m/s, and 46 m/s, respectively. The recorded RBVmean value shows the unmistakable potential of the OKV. By comparison, average values for a similarly sized SJM Regent valve exceed 100 m/s and fall between 10 â 30 m/s for bioprosthetic valves. The in-vitro results for both the closing volume and RBVmean of the OKV prototype demonstrate the potential of the novel BMHV design. Conclusions and Future Works 65 Conclusions and Future Works In this thesis, a brief overview of valvular heart disease and the options that are available to afflicted patients was provided. Heart valve prosthetics have had a long history of development and evolution since their invention, but modern research has neglected mechanical heart valves for many years. The research summarized in this thesis was conducted as an effort to examine the primary design limitations afflicting mechanical heart valves and to design and evaluate a novel valve that addressed the identified limitations. This work culminated in the design of the Okanagan Valve which is proposed to serve as a basis for the next generation of bileaflet mechanical heart valves. In-silico and in-vitro analysis techniques were leveraged to evaluate the performance of the proposed valve and to compare it to the industry gold standard. This chapter serves to conclude this thesis by discussing the primary findings, the limitations of the design and evaluation work, and possible avenues for future works that build upon these conclusions. 6.1 Discussion In summary, this thesis identified the perceived limitations of BMHVs, developed a novel design architecture that addressed these limitations, and evaluated that design numerically and experimentally. In Chapter 2, the design limitations of bileaflet mechanical heart valves were discussed as their hemodynamics, durability, and biocompatibility. These limitations were utilized as the basis for the Okanagan Valveâs design in Chapter 3. â¢ The hemodynamics of BMHVs were evaluated in terms of shear stresses, flow stagnation, regurgitation, effective orifice area, and cavitation. â¢ The durability of BMHVs known failure modes were evaluated. â¢ The biocompatibility of implanted BMHVs were evaluated in terms of the materials used to construct the valves and the suture rings that hold them in place. Conclusions and Future Works 66 In Chapter 3, the design of the Okanagan Valve was proposed, and how its design addresses the limitations identified in Chapter 2 was discussed. â¢ The design of the Okanagan Valve was broken down into the housing, leaflets, and hinge mechanism. The design was shown and described through engineering views and renders. â¢ An overview of the critical design elements of the Okanagan Valve and how they address the identified limitations of modern BMHVs was given. In Chapter 4, the methods utilized to evaluate the performance of the Okanagan Valve were discussed. â¢ A numerical analysis was performed to simulate regurgitant flow from the aortic root into the left ventricle and leaflet behavior in those conditions. â¢ A quick flow simulation was performed to estimate and visualize flow patterns in standard BMHVs compared to the Okanagan Valve. â¢ In-vitro analysis of the Okanagan Valve was performed to experimentally assess its performance. In-vitro analysis was preceded by steady-state testing of prototypes to ensure their adequacy. In Chapter 5, the performance of the Okanagan Valve was evaluated leveraging the methods discussed in Chapter 4. â¢ The numerical analysis found the novel leaflet shape produced favourable results compared to the standard leaflet shape of the SJM valve. Introducing ovality to the leaflet shape affected the performance and it was found that a major diameter of 27 mm and a minor diameter of 23 mm performed optimally. â¢ The optimal ovality was numerically found to have a closing volume of 5.3 cc, closing time of 17 ms, maximum blood velocity of 1.7 m/s, and maximum leaflet tip velocity of 1.6 m/s. Conclusions and Future Works 67 â¢ Steady-state leakage tests were performed with a pressure of 120/80 mmHg and a saline solution. The Okanagan Valve aluminum prototype was found to have an adequate geometric leakage area of 0.04 cm2. â¢ In-vitro experimental analysis of the aluminum prototype in a pulse duplicator found it to have a closing volume of approximately 5 cc, a closing time of 35 ms, and a mean regional backflow velocity of 44 m/s. As recorded, the experimental results are reminiscent of a bioprosthetic heart valve, unlike those of a traditional BMHV. 6.2 Limitations This thesis was a preliminary study into the kinematics of the Okanagan Valve and its viability a novel BMHV architecture. Some noted limitations of the work completed are discussed below. 6.2.1 Prototype Manufacturing Producing a prototype with adequate detail and material properties was an issue throughout the study. The intricate geometries of the OKV are prohibitively difficult to produce with conventional machining techniques or additive manufacturing technologies. While the detailed geometries could be reproduced with many 3D printing techniques, the materials used lacked the requisite properties for in-vitro analysis. The aluminum prototype manufactured using the DMLS process was sufficiently strong, however, the poor surface finish and resolution of the print hindered its performance. 6.2.2 Prototype Material Due to cost and manufacturing limitations, prototypes could not be manufactured out of pyrolytic carbon, which would be the final material. The material properties and surface finish of pyrolytic carbon alter the possible allowances for a prototype and would affect the experimental results by an unknown factor. 6.2.3 Hinge Geometry Reproduction Due to limitations in prototype manufacturing, an accurate butterfly hinge could not be reproduced for in-vitro analysis. Pinned pivot mechanisms were substituted to allow for leaflet actuation in the tested prototypes. It is important to model the hinge flow mechanics Conclusions and Future Works 68 of an MHV accurately as the hinges are a common point of failure. The mechanical allowances required for a pinned pivot differ from those of a butterfly hinge and could have introduced error into our experimental results. 6.2.4 Minimal In-Vitro Testing Ability The in-vitro analysis was conducted at an external lab due to the lack of in-vitro testing equipment in our lab. This hindered our ability to adjust and iterate on the OKV and run successive tests. Our conclusions were based on data from five systolic cycles due to time limitations with the equipment. Analysis at varying heart rates and cardiac outputs could not be conducted. 6.3 Future Works The limitations of this thesis are potential avenues for future research to build upon the work that was completed. Some of the identified directions for future works are discussed below. 6.3.1 Housing Inlet/Outlet Optimization It would be beneficial for the housing design to match the native anatomy closer than the approximated saddle shape. Flaring the inlet and outlet to fit against the native annulus would improve valve placement, mitigate paravalvular leakage, and reduce fluid resistance through the valve. 6.3.2 Optimizing Ovality Ovality was implemented into the OKV design with the rotational axis set along the major diameter. Further research should be conducted on finding the optimum ratio between the major and minor diameters to improve the hemodynamics of the OKV. It would also be interesting to evaluate ovality with the axis of rotation set along the minor diameter or even offset from it. 6.3.3 Hinge Design Accurate analysis of hinge wash was not conducted in this study and is a critical component of MHV design. Experimental data to corroborate the theoretical improvements would benefit the credibility of the OKV design. It would be beneficial to examine the novel socket-on-leaflet approach of the OKV and evaluate the ideal geometry to promote wash. It would Conclusions and Future Works 69 be valuable to determine how altering the size of the hinge would affect its hemodynamics. Furthermore, the width of the hingeâs male component is not limited by the thickness of the leaflet in the OKV design allowing for a larger or smaller hinge to be incorporated. 6.3.4 Leaflet Curvature Optimization Further in-vitro and numerical analysis of the leaflet curvature could be conducted to determine its effect on flow through the valve. It would be beneficial for the design to evaluate additional radii for the curvature and measure its effect on the OKVâs hemodynamics. 6.3.5 Rapid Prototype Manufacturing and Local In-Vitro Analysis Low-force stereolithographic 3D printing technology is potentially capable of producing valve prototypes both detailed enough for accurate hinge geometry and from materials adequate for in-vitro analysis. The acquisition of in-vitro testing equipment for our lab would be immensely beneficial to the further development of the OKV and would enable greater confidence in the experimental results through replication. The combined addition of both technologies to our lab would enable a rapid iterative approach to the design as well as experimental confirmation of hemodynamic performance for each iteration. 6.3.6 Addressing Endocarditis Endocarditis is a serious complication of SAVR that can occur following surgery. Our study did not address contributing factors to the onset of endocarditis or ways to mitigate infection. It would be a great asset for a next-generation BMHV to address this common contributing factor to patient mortality. 6.3.7 Pediatric OKV It would be valuable to evaluate the novel BMHV architecture proposed by the OKV for pediatric applications. Its hemodynamic performance, as suggested by this study, indicates that it may perform favorably at diameters <20 mm and in younger patients who have a greater bleeding risk due to injuries and anticoagulant therapy. The OKV may be an excellent candidate for a pediatric MHV to treat congenital heart valve defects which are currently a relatively unmet need. A previous study conducted in our lab showed that current BMHVs perform poorly at elevated heart rates which are a defining characteristic of youth. It would Conclusions and Future Works 70 be interesting to further study the performance of the OKV at physiological conditions representative of youth to determine its applicability as a solution for younger patients. The human heart is a muscular organ responsible for moving blood around the body and through the lungs. Contained within the heart are valves which convert the contractions of the heart into pumping forces. Valvular heart disease afflicts millions of individuals worldwide with potentially fatal consequences if allowed to progress unchecked. As prosthetic heart valves today suffer from a wide variety of side effects and performance limitations, the development of a true, permanent replacement for diseased native heart valves is the holy grail of heart valve research. The research compiled in this thesis suggests that the unique design of the OKV has the potential to shift the current landscape of bileaflet mechanical heart valves and improve the prognosis of individuals diagnosed with valvular heart disease.Bibliography 71 Bibliography [1] B. Iung and A. Vahanian, âEpidemiology of acquired valvular heart disease,â Canadian Journal of Cardiology, vol. 30, no. 9. pp. 962â970, Sep-2014. [2] U.S. Department of Health & Human Services and National Heart Lung and Blood Institute, âHeart Valve Disease,â 2015. [Online]. Available: https://www.nhlbi.nih.gov/health-topics/heart-valve-disease. [Accessed: 22-Oct-2018]. [3] J. Rhead (Intermountain Healthcare), âAortic Valve Stenosis | Heart Care.â [Online]. Available: https://intermountainhealthcare.org/services/heart-care/conditions/aortic-valve-stenosis/. [Accessed: 17-Nov-2019]. [4] C. M. Otto, âEvaluation and Management of Chronic Mitral Regurgitation,â N. Engl. J. Med., vol. 345, no. 10, pp. 740â746, Sep. 2001. [5] K. Maganti, V. H. Rigolin, M. E. Sarano, and R. O. Bonow, âValvular heart disease: diagnosis and management.,â Mayo Clin. Proc., vol. 85, no. 5, pp. 483â500, May 2010. [6] J. R. Carapetis et al., âAcute rheumatic fever and rheumatic heart disease.,â Nat. Rev. Dis. Prim., vol. 2, p. 15084, 2016. [7] A. P. Yoganathan, Z. He, and S. Casey Jones, âFluid Mechanics of Heart Valves,â Annu. Rev. Biomed. Eng., vol. 6, no. 1, pp. 331â362, Aug. 2004. [8] B. R. Lindman et al., âCalcific aortic stenosis,â Nat. Rev. Dis. Prim., vol. 2, p. 16006, 2016. [9] H. Mohammadi and G. Fradet, âProsthetic Aortic Heart Valves,â Cardiovasc. Syst., vol. 5, no. 1, p. 2, 2017. [10] J. Ennker, A. Albert, and I. C. Ennker, âStentless aortic valves. Current aspects.,â HSR Proc. Intensive Care Cardiovasc. Anesth., vol. 4, no. 2, pp. 77â82, 2012. [11] A. Schaefer et al., âStentless vs. stented bioprosthesis for aortic valve replacement: A case matched comparison of long-term follow-up and subgroup analysis of patients with native valve endocarditis,â PLoS One, vol. 13, no. 1, p. e0191171, Jan. 2018. [12] D. Mylotte, R. Lange, G. Martucci, and N. Piazza, âTranscatheter heart valve implantation for failing surgical bioprostheses: technical considerations and evidence for valve-in-valve procedures.,â Heart, vol. 99, no. 13, pp. 960â7, Jul. 2013. [13] D. Y. Tam, T. X. Vo, H. C. Wijeysundera, D. Dvir, J. O. Friedrich, and S. E. Fremes, âTranscatheter valve-in-valve versus redo surgical aortic valve replacement for the treatment of degenerated bioprosthetic aortic valve: A systematic review and meta-analysis,â Catheter. Cardiovasc. Interv., Jul. 2018. [14] J. K. Forrest, âTranscatheter aortic valve replacement: design, clinical application, and Bibliography 72 future challenges.,â Yale J. Biol. Med., vol. 85, no. 2, pp. 239â47, Jun. 2012. [15] L. N. Scotten and R. Siegel, âAre anticoagulant independent mechanical valves within reach-fast prototype fabrication and in vitro testing of innovative bi-leaflet valve models.,â Ann. Transl. Med., 2015. [16] N. C. Chan, J. I. Weitz, and J. W. Eikelboom, âAnticoagulation for Mechanical Heart Valves,â Arterioscler. Thromb. Vasc. Biol., vol. 37, no. 5, pp. 743â745, May 2017. [17] M. T. Crim, S. Shore, S. K. Lee, and B. J. Wells, âAnticoagulation for cardiac prosthetic devices: Prosthetic heart valves, left ventricular assist devices, and septal closure devices,â in Anticoagulation Therapy, Springer International Publishing, 2018, pp. 253â295. [18] T. W. Secomb, âHemodynamics.,â Compr. Physiol., vol. 6, no. 2, pp. 975â1003, Mar. 2016. [19] G. D. Dangas, J. I. Weitz, G. Giustino, R. Makkar, and R. Mehran, âProsthetic Heart Valve Thrombosis,â Journal of the American College of Cardiology, vol. 68, no. 24. Elsevier USA, pp. 2670â2689, 20-Dec-2016. [20] A. S. Wolberg, M. M. Aleman, K. Leiderman, and K. R. Machlus, âProcoagulant activity in hemostasis and thrombosis: Virchowâs triad revisited.,â Anesth. Analg., vol. 114, no. 2, pp. 275â85, Feb. 2012. [21] W. Y. Lim, G. Lloyd, and S. Bhattacharyya, âMechanical and surgical bioprosthetic valve thrombosis,â Heart, vol. 103, no. 24. BMJ Publishing Group, pp. 1934â1941, 01-Dec-2017. [22] J. Ding et al., âQuantification of Shear-Induced Platelet Activation: High Shear Stresses for Short Exposure Time,â Artif. Organs, vol. 39, no. 7, pp. 576â583, Jul. 2015. [23] B. H. Jun, N. Saikrishnan, S. Arjunon, B. M. Yun, and A. P. Yoganathan, âEffect of hinge gap width of a St. Jude medical bileaflet mechanical heart valve on blood damage potential-an in vitro micro particle image velocimetry study,â J. Biomech. Eng., vol. 136, no. 9, 2014. [24] G. Girdhar and D. Bluestein, âBiological effects of dynamic shear stress in cardiovascular pathologies and devices.,â Expert Rev. Med. Devices, vol. 5, no. 2, pp. 167â81, Mar. 2008. [25] Y. Alemu and D. Bluestein, âFlow-induced Platelet Activation and Damage Accumulation in a Mechanical Heart Valve: Numerical Studies,â Artif. Organs, vol. 31, no. 9, pp. 677â688, Sep. 2007. [26] D. L. Bark et al., âHemodynamic Performance and Thrombogenic Properties of a Superhydrophobic Bileaflet Mechanical Heart Valve,â Ann. Biomed. Eng., vol. 45, no. 2, pp. 452â463, Feb. 2017. [27] P. R. Hoskins, P. V. Lawford, and B. J. Doyle, Cardiovascular Biomechanics. Springer Bibliography 73 International Publishing, 2017. [28] V. Govindarajan, H. S. Udaykumar, and K. B. Chandran, âFLOW DYNAMIC COMPARISON BETWEEN RECESSED HINGE AND OPEN PIVOT BI-LEAFLET HEART VALVE DESIGNS.â [29] B. M. Yun et al., âA numerical investigation of blood damage in the hinge area of aortic bileaflet mechanical heart valves during the leakage phase,â Ann. Biomed. Eng., vol. 40, no. 7, pp. 1468â1485, Jul. 2012. [30] P. C. Lu, J. S. Liu, R. H. Huang, C. W. Lo, H. C. Lai, and N. H. C. Hwang, âThe closing behavior of mechanical aortic heart valve prostheses,â ASAIO J., vol. 50, no. 4, pp. 294â300, 2004. [31] D. H. Lee et al., âThe measurement of opening angle and orifice area of a bileaflet mechanical valve using multidetector computed tomography,â Korean Circ. J., vol. 39, no. 4, pp. 157â162, Apr. 2009. [32] H. Mohammadi and G. Fradet, âOval housing for the St. Jude Medical bileaflet mechanical heart valve,â Proc. Inst. Mech. Eng. Part H J. Eng. Med., vol. 231, no. 10, pp. 982â986, Oct. 2017. [33] E. Klusak, A. Bellofiore, S. Loughnane, and N. J. Quinlan, âHigh-resolution measurements of velocity and shear stress in leakage jets from bileaflet mechanical heart valve hinge models,â J. Biomech. Eng., vol. 137, no. 10, Oct. 2015. [34] M. Hedayat, H. Asgharzadeh, and I. Borazjani, âPlatelet activation of mechanical versus bioprosthetic heart valves during systole,â J. Biomech., vol. 56, pp. 111â116, May 2017. [35] H. A. Simon, L. Ge, F. Sotiropoulos, and A. P. Yoganathan, âNumerical Investigation of the Performance of Three Hinge Designs of Bileaflet Mechanical Heart Valves,â Ann. Biomed. Eng., vol. 38, no. 11, pp. 3295â3310, Nov. 2010. [36] F. Sotiropoulos, T. B. Le, and A. Gilmanov, âFluid Mechanics of Heart Valves and Their Replacements,â Annu. Rev. Fluid Mech., vol. 48, no. 1, pp. 259â283, Jan. 2016. [37] S. J. Head, M. Ãelik, and A. P. Kappetein, âMechanical versus bioprosthetic aortic valve replacement,â European Heart Journal, vol. 38, no. 28. Oxford University Press, pp. 2183â2191, 21-Jul-2017. [38] P. Pibarot and J. G. Dumesnil, âProsthesis-patient mismatch: Definition, clinical impact, and prevention,â Heart, vol. 92, no. 8. pp. 1022â1029, Aug-2006. [39] S. J. Head et al., âThe impact of prosthesis-patient mismatch on long-term survival after aortic valve replacement: a systematic review and meta-analysis of 34 observational studies comprising 27 186 patients with 133 141 patient-years.â [40] P. B. Bertrand et al., âEffective Orifice Area during Exercise in Bileaflet Mechanical Valve Prostheses,â J. Am. Soc. Echocardiogr., vol. 30, no. 4, 2017. [41] L. J. Castro, J. M. Arcidi, A. L. Fisher, and V. A. Gaudiani, âRoutine enlargement of the Bibliography 74 small aortic root: A preventive strategy to minimize mismatch,â in Annals of Thoracic Surgery, 2002, vol. 74, no. 1, pp. 31â36. [42] H. Lee, Y. Taenaka, and S. Kitamura, âMechanism for cavitation in the mechanical heart valve with an artificial heart: Nuclei and viscosity dependence,â Artif. Organs, vol. 29, no. 1, pp. 41â46, Jan. 2005. [43] D. Bluestein, S. Einav, and N. H. C. Hwang, âA squeeze flow phenomenon at the closing of a bileaflet mechanical heart valve prosthesis,â J. Biomech., vol. 27, no. 11, pp. 1369â1378, 1994. [44] H. Lee, A. Homma, E. Tatsumi, and Y. Taenaka, âObservation of cavitation pits on mechanical heart valve surfaces in an artificial heart used in in vitro testing,â J. Artif. Organs, vol. 13, no. 1, pp. 17â23, Apr. 2010. [45] H. Lee, E. Tatsumi, A. Homma, T. Tsukiya, and Y. Taenaka, âMechanism for cavitation of monoleaflet and bileaflet valves in an artificial heart,â J. Artif. Organs, vol. 9, no. 3, pp. 154â160, Sep. 2006. [46] S. L. Poliachik, W. L. Chandler, R. J. Ollos, M. R. Bailey, and L. A. Crum, âThe relation between cavitation and platelet aggregation during exposure to high-intensity focused ultrasound,â Ultrasound Med. Biol., vol. 30, no. 2, pp. 261â269, Feb. 2004. [47] P. Johansen, âMechanical heart valve cavitation,â Expert Review of Medical Devices, vol. 1, no. 1. pp. 95â104, Sep-2004. [48] M. Barak and Y. Katz, âMicrobubbles: Pathophysiology and clinical implications,â Chest, vol. 128, no. 4, pp. 2918â2932, 2005. [49] R. A. Rodriguez, M. Ruel, M. Labrosse, and T. Mesana, âTranscranial Doppler and acoustic pressure fluctuations for the assessment of cavitation and thromboembolism in patients with mechanical heart valves,â Interact. Cardiovasc. Thorac. Surg., vol. 7, no. 2, pp. 179â183, 2007. [50] H. Zhang et al., âPivoting system fracture in a bileaflet mechanical valve: A case report,â J. Biomed. Mater. Res. Part B Appl. Biomater., vol. 90B, no. 2, pp. 952â961, Aug. 2009. [51] Y. Misawa, âValve-related complications after mechanical heart valve implantation,â Surgery Today, vol. 45, no. 10. Springer Tokyo, pp. 1205â1209, 14-Oct-2015. [52] S. P. Collison and Y. K. Mishra, âLeaflet escape in a modern mechanical heart valve: Structure and cavitation,â Heart, vol. 96, no. 12, p. 994, 2010. [53] E. M. Baudet et al., âLong-term results of valve replacement with the St. Jude Medical prosthesis,â J. Thorac. Cardiovasc. Surg., vol. 109, no. 5, pp. 858â870, 1995. [54] Y. Misawa, K. Fuse, T. Saito, H. Konishi, and S. I. Oki, âFourteen year experience with the omnicarbon prosthetic heart valve,â ASAIO J., vol. 47, no. 6, pp. 677â682, 2001. [55] Y. Misawa et al., âClinical experience with the Bicarbon heart valve prosthesis,â J. Bibliography 75 Cardiothorac. Surg., vol. 2, no. 1, p. 8, Dec. 2007. [56] R. Biran and D. Pond, âHeparin coatings for improving blood compatibility of medical devices,â Advanced Drug Delivery Reviews, vol. 112. Elsevier B.V., pp. 12â23, 01-Mar-2017. [57] V. L. Gott and R. L. Daggett, âSerendipity and the development of heparin and carbon surfaces,â in Annals of Thoracic Surgery, 1999, vol. 68, no. 3 SUPPL. [58] V. L. Gott, D. E. Alejo, and D. E. Cameron, âMechanical Heart Valves: 50 Years of Evolution,â in Annals of Thoracic Surgery, 2003, vol. 76, no. 6. [59] S. W. Jordan and E. L. Chaikof, âNovel thromboresistant materials,â J. Vasc. Surg., vol. 45, no. 6 SUPPL., pp. A104âA115, 2007. [60] H. Zeng, W. Yin, G. Catausan, N. Moldovan, and J. Carlisle, âUltrananocrystalline diamond integration with pyrolytic carbon components of mechanical heart valves,â Diam. Relat. Mater., vol. 61, pp. 97â101, Jan. 2016. [61] V. Jokinen, E. Kankuri, S. Hoshian, S. Franssila, and R. H. A. Ras, âSuperhydrophobic Blood-Repellent Surfaces,â Adv. Mater., vol. 30, no. 24, p. 1705104, Jun. 2018. [62] C. A. Mestres, A. Igual, and M. Murtra, âClinical performance of the Omniscience prosthetic heart valve with Dacron sewing ring,â Thorac. Cardiovasc. Surg., vol. 33, no. 5, pp. 296â299, 1985. [63] B. L. Illingworth, K. Tweden, R. F. Schroeder, and J. D. Cameron, âIn vivo efficacy of silver-coated (Silzone(TM)) infection-resistant polyester fabric against a biofilm-producing bacteria, Staphylococcus epidermidis,â J. Heart Valve Dis., vol. 7, no. 5, pp. 524â530, Sep. 1998. [64] R. G. Seipelt et al., âThe St. Jude âSilzoneâ valve: Midterm results in treatment of active endocarditis,â Ann. Thorac. Surg., vol. 72, no. 3, pp. 758â762, 2001. [65] G. Grunkemeier, âThe Silzone effect: how to reconcile contradictory reports?,â Eur. J. Cardio-Thoracic Surg., vol. 25, no. 3, pp. 371â375, Mar. 2004. [66] A. Ionescu, N. Payne, A. G. Fraser, J. Giddings, G. L. Grunkemeier, and E. G. Butchart, âIncidence of embolism and paravalvar leak after St Jude Silzone valve implantation: Experience from the Cardiff Embolic Risk Factor Study,â Heart, vol. 89, no. 9, pp. 1055â1061, Sep. 2003. [67] J. Butany et al., âPathologic Analysis of 19 Heart Valves With Silver-Coated Sewing Rings,â J. Card. Surg., vol. 21, no. 6, pp. 530â538, Nov. 2006. [68] M. Cimbollek, B. Nies, R. Wenz, and J. Kreuter, âAntibiotic-impregnated heart valve sewing rings for treatment and prophylaxis of bacterial endocarditis.,â Antimicrob. Agents Chemother., vol. 40, no. 6, pp. 1432â7, Jun. 1996. [69] B. G. French, K. Wilson, M. Wong, S. Smith, and M. F. OâBrien, âRifampicin antibiotic Bibliography 76 impregnation of the St. Jude Medical mechanical valve sewing ring: A weapon against endocarditis,â J. Thorac. Cardiovasc. Surg., vol. 112, no. 2, pp. 248â252, 1996. [70] M. Karck, E. Siclari, H. Wahlig, U. Sperling, C. Schmid, and A. Haverich, âPretreatment of prosthetic valve sewing-ring with the antibiotic/fibrin sealant compound as a prophylactic tool against prosthetic valve endocarditis,â Eur. J. Cardio-thoracic Surg., vol. 4, no. 3, pp. 142â146, 1990. [71] Y. H. Cho, K. Sung, W. S. Kim, J. H. Kim, S. M. Kim, and P. W. Park, âMechanical surface area of prosthetic heart valve: Adverse clinical impact of large mechanical valve in mitral position,â ASAIO J., vol. 64, no. 6, pp. 779â784, 2018. [72] S. K. Kadhim, M. S. Nasif, H. H. Al-Kayiem, and R. Al-Waked, âComputational fluid dynamics simulation of blood flow profile and shear stresses in bileaflet mechanical heart valve by using monolithic approach,â Simulation, vol. 94, no. 2, 2018. [73] ViVitro Labs, âSuperPump Pulsatile Pump.â [Online]. Available: https://vivitrolabs.com/product/superpump-pulsatile-blood-pump/. [Accessed: 04-Jul-2019]. [74] ViVitro Labs, âInterview with Larry Scotten and VSIâs prototype Leonardo | Vivitro Labs.â [Online]. Available: https://vivitrolabs.com/larry-scotten-and-leonardo-vsi/. [Accessed: 09-Jul-2019]. Appendices 77 Appendices Appendix A: FFF Prototypes The following sections contain additional images depicting different views of 3D printed prototypes of OKV components and gaskets. Figure A.1: Upscaled OKV leaflet with multiple print defects. (Left) The exterior surface of the leaflet. (Right) The side view of the leaflet displaying the hinge geometry. Figure A.2: The bottom view of an early full-scale OKV prototype, highlighting the pinned hinges. Appendices 78 Figure A.3: Two views of a full-scale OKV prototype; featuring improved print quality, functional hinges, and straight leaflets. Figure A.4: Two views of a full-scale OKV prototype; featuring a revised leaflet tip geometry, flat housing outlet, and a more aggressive inlet cutout. Appendices 79 Figure A.5: Two views of a near-final revision full-scale OKV prototype. Printed using an FFF 3D printer, finished by hand, and painted to represent the appearance of a realistic BMHV. Figure A.6: Two views of a full-scale OKV housing prototype printed from PLA without a suture ring. Appendices 80 Figure A.7: TPU Gasket Layer Separation (Closeup View) Appendices 81 Appendix B: Flow Simulation The following sections contain all images generated for the flow simulations conducted on the OKV and BMHV, as well as views of the simulated aortic root. These images are for visual reference only and should not be used to measure the performance of either depicted valve. B.1 Side Views of Both Designs Figure A.8: OKV (left) and BMHV (right) Side Profile at 16.6% Open Figure A.9: OKV (left) and BMHV (right) Side Profile at 33.2% Open Appendices 82 Figure A.10: OKV (left) and BMHV (right) Side Profile at 49.8% Open Figure A.11: OKV (left) and BMHV (right) Side Profile at 66.4% Open Figure A.12: OKV (left) and BMHV (right) Side Profile at 83.0% Open Appendices 83 Figure A.13: OKV (left) and BMHV (right) Side Profile at 100.0% Open B.2 Angled Views of Both Designs Figure A.14: OKV (left) and BMHV (right) at 16.6% Open Figure A.15: OKV (left) and BMHV (right) at 33.2% Open Appendices 84 Figure A.16: OKV (left) and BMHV (right) at 49.8% Open Figure A.17: OKV (left) and BMHV (right) at 66.4% Open Figure A.18: OKV (left) and BMHV (right) at 83.0% Open Appendices 85 Figure A.19: OKV (left) and BMHV (right) at 100.0% Open Appendices 86 B.3 Views of the Simulated Aortic Root The following views show the simulated aorta that was designed and used for the simulation. Figure A.20: Exterior Render of the Simulated Aorta (Front View) Figure A.21: Exterior Render of the Simulated Aorta (Side View) Appendices 87 Appendix C: Okanagan Valve Renders The following section contains images of the Okanagan Valve rendered using SolidWorks. The Okanagan Valve in the open position, closed position, the housing, and a single leaflet are shown in five different orientations. C.1 Side view Figure A.22: Four rendered side views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. Appendices 88 C.2 Front view Figure A.23: Four rendered front views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. Appendices 89 C.3 Top view Figure A.24: Four rendered top views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. Appendices 90 C.4 Bottom view Figure A.25: Four rendered bottom views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. Appendices 91 C.5 Isometric view Figure A.26: Four rendered bottom views of the Okanagan Valve: (A) closed position, (B) open position, (C) the housing, (D) a leaflet. Appendices 92 Appendix D: Engineering Drawings The following sections contain engineering drawings of the Okanagan Valve produced in SolidWorks. PAGE: 93OKV LEAFLETPROPRIETARY AND CONFIDENTIALTHE INFORMATION CONTAINED IN THIS DRAWING IS THE SOLE PROPERTY OF THE HVPL. ANY REPRODUCTION IN PART OR AS A WHOLE WITHOUT THE WRITTEN PERMISSION OF THE HVPL IS PROHIBITED.DIMENSIONS ARE IN MMAppendix D.1 DO NOT SCALE DRAWINGMATERIALS: Pyrolitic Carbon Coated Graphite and DacronDWG. NO.SCALE:3:1A AB B2211SOLIDWORKS Educational Product. For Instructional Use Only.AABBSECTION A-ASECTION B-BPAGE: 94OKV HOUSINGPROPRIETARY AND CONFIDENTIALTHE INFORMATION CONTAINED IN THIS DRAWING IS THE SOLE PROPERTY OF THE HVPL. ANY REPRODUCTION IN PART OR AS A WHOLE WITHOUT THE WRITTEN PERMISSION OF THE HVPL IS PROHIBITED.DIMENSIONS ARE IN MMAppendix D.2 DO NOT SCALE DRAWINGMATERIALS: Pyrolitic Carbon Coated Graphite and DacronDWG. NO.SCALE:2:1A AB B2211SOLIDWORKS Educational Product. For Instructional Use Only.
Thesis/Dissertation
eng
Mechanical Engineering
Vancouver : University of British Columbia Library
Attribution-NonCommercial-NoDerivatives 4.0 International
Graduate
Okanagan valve : the next generation of mechanical heart valves
Text