design, development and evaluation of centrifugal ventricular

Design, Development and Evaluation of Centrifugal Ventricular Assist Devices

Submitted by

Daniel L. Timms

Bachelor of Mechanical Engineering (Honours)

Queensland University of Technology

A thesis submitted in total fulfilment

of the requirements for the degree of

Doctor of Philosophy

School of Mechanical, Manufacturing and Medical Engineering

Faculty of Built Environment and Engineering

Queensland University of Technology

Brisbane, Queensland, Australia

2005

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Keywords Left- Right- Bi-Ventricular Assist Device, Mock Circulation Loop, Centrifugal Pump

Design, Impeller Hydraulic Force and Performance.

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Abstract Heart disease is the developed world's biggest killer, and the shortage of donor hearts

has accelerated the development of mechanical alternatives.

Scientists, engineers and clinicians have attempted to replicate the human heart with

a mechanical device for over 50 years. Although a number of pulsating devices have

been developed, and in some cases worked briefly, they have invariably failed to

match the success of heart transplantation.

In an attempt to produce a suitable alternative, current research is focused on devices

that do not replace the heart; but rather work along side it to assist its function. Many

of these devices help the failing left ventricle; however some patients require the

additional implantation of a second device to assist a failing right ventricle. This

increases implantation time and associated risk, and because of the size of the current

devices, reduces the access of smaller patients to this vital technology.

The overall thesis objective focuses on the progressive design, development and

preliminary evaluation of two novel centrifugal type ventricular assist devices, a bi-

left ventricular device (Bi-LVAD) and a single bi-ventricular assist device (Bi-

VAD). The devices have the respective capability to assist either the left ventricle, or

both ventricles of a failing heart. The current concept for each VAD employs both

magnetic and hydrodynamic suspension techniques to float a rotating double

impeller, a technique that aims to reduce blood damage and component wear, two of

the major problems encountered with current generation devices.

Each VAD design was developed by conducting experimentation and drawing

conclusions from a variety of engineering research fields, such as flow visualization,

rotary pump design and testing, fluid dynamics, hemodynamics and heart failure, and

magnetic motor bearing design.

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In order to evaluate pump prototype designs, it was necessary to design and develop

a novel pulsatile systemic and pulmonary mock circulation loop capable of

reproducing the hemodynamics of heart failure in the systemic and pulmonary

circuits. The investigation then specifically examined the static hydraulic forces on

the impeller of a centrifugal blood pump during operation in this mock circulation

loop. The recorded magnitude and direction of radial and axial thrust then influenced

the selection of magnetic and hydrodynamic bearing configurations to minimise

impeller touchdown in the intended hemodynamic environment. This research

required the development of correctly designed impeller (semi-open/closed) and

volute (single, double, circular) components for each ventricular assist application

and a unique test facility to isolate impeller hydraulic forces in addition to the mock

circulation loop.

The proposed Bi-LVAD incorporates symmetrical blade designs on each side of the

double sided impeller. The device assists the function of the left ventricle only with

symmetrical axial pressure distribution and elimination of stagnant regions beneath

the impeller. These features improve axial touchdown capacity and reduce thrombus

formation respectively. The proposed Bi-VAD incorporates different blade designs

on each side of the double impeller to augment the function of both the left and right

cardiac chambers. The design has the additional potential to act as a total artificial

heart (TAH). To date there is no Bi-VAD/TAH system available that incorporates an

LVAD and RVAD in one rotary pump.

Successful development of each innovative VAD will provide an alternative to heart

transplantation, potentially saving lives of many terminal heart patients each year. No

longer would heart transplant candidates need to wait for the untimely death of a

donor to provide a suitable heart. Instead, this new generation device would be

available immediately, and be almost universally compatible with all patients. It has

the potential to dramatically increase a patient’s expected lifetime, and to deliver

them a higher quality of life.

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Table of Contents Design, Development and Evaluation of Centrifugal Ventricular Assist Devices.. i Keywords ................................................................................................................... ii Abstract .................................................................................................................. iii Table of Contents .........................................................................................................v List of Figures............................................................................................................. ix List of Tables ............................................................................................................. xii Nomenclature ........................................................................................................... xiv Statement of Authorship............................................................................................xv Acknowledgements................................................................................................... xvi Chapter 1 Introduction.......................................................................................... 1-1

1.1 Significance................................................................................................. 1-5 1.2 Objectives.................................................................................................... 1-6 1.3 Aims ............................................................................................................ 1-7 1.4 Thesis Outline ............................................................................................. 1-8 1.5 Thesis Outcomes and Contributions ......................................................... 1-10 1.6 Publications ............................................................................................... 1-11

Chapter 2 Literature Review................................................................................. 2-1

2.1 Heart Failure and Treatment Options.......................................................... 2-2 2.2 Ventricular Assist Device Design Considerations ...................................... 2-6 2.3 Flow Visualisation .................................................................................... 2-12 2.4 Right Heart Failure and Bi-Ventricular Support....................................... 2-14 2.5 Current Mechanical Assist Devices .......................................................... 2-17 2.6 Design, Performance and Force Characteristics ....................................... 2-22 2.7 Magnetic Bearings for Rotary Blood Pumps ............................................ 2-36 2.8 Review Summary ...................................................................................... 2-45 2.9 Conclusion ................................................................................................ 2-49

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Chapter 3 Mock Circulation Loop........................................................................ 3-1 3.1 Introduction ................................................................................................. 3-2 3.2 Background - Description of Physiological Parameters ............................. 3-2

3.2.1 Circulatory System ...................................................................................................3-2 3.2.2 Blood Volume and Pressure Distribution .................................................................3-3 3.2.3 Heart Functionality...................................................................................................3-4 3.2.4 Vascular Hemodynamics..........................................................................................3-6 3.2.5 Physiological Pressures ............................................................................................3-7 3.2.6 Compliance............................................................................................................. 3-11 3.2.7 Resistance............................................................................................................... 3-13 3.2.8 Cardiac Output ....................................................................................................... 3-16 3.2.9 Cardiac / Vascular Coupling................................................................................... 3-19 3.2.10 Autoregulation........................................................................................................ 3-24

3.3 Literature Review – Prototype VAD Testing............................................ 3-25 3.3.1 Basic Mock Circulation Loops ............................................................................... 3-25 3.3.2 Advanced Mock Circulation Loops........................................................................ 3-26 3.3.3 Current Mock Loops............................................................................................... 3-26 3.3.4 In-Vivo Reproduction of Heart Failure .................................................................. 3-34

3.4 Simulation Model of Circulatory System.................................................. 3-35 3.4.1 Simulation Circulation System Configurations ...................................................... 3-35 3.4.2 Results .................................................................................................................... 3-37

3.5 Experimental Mock Circulation Loop....................................................... 3-39 3.5.1 Design Criteria ....................................................................................................... 3-39 3.5.2 Description of Mock Circulation Loop................................................................... 3-39 3.5.3 Experimental Procedure ......................................................................................... 3-52 3.5.4 Results .................................................................................................................... 3-55

3.6 Discussion ................................................................................................. 3-57 3.6.1 Heart Functionality................................................................................................. 3-57 3.6.2 Pressure Distribution .............................................................................................. 3-60 3.6.3 Vasculature............................................................................................................. 3-63 3.6.4 Perfusion................................................................................................................. 3-65 3.6.5 Cardiovascular Interaction and Autoregulation...................................................... 3-66

3.7 Conclusion................................................................................................. 3-67

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Chapter 4 Centrifugal VAD Design and Development....................................... 4-1 4.1 Introduction................................................................................................. 4-2 4.2 Background ................................................................................................. 4-4

4.2.1 Centrifugal Pump Design .........................................................................................4-4 4.2.2 Hydraulic Performance...........................................................................................4-12 4.2.3 Hydraulic Force ......................................................................................................4-15

4.3 Design Procedure and Parameter Calculation........................................... 4-32 4.3.1 Design Conditions and Fluid Properties .................................................................4-33 4.3.2 Pump Constants from Similitude............................................................................4-34

4.4 Design Detail............................................................................................. 4-36 4.4.1 LVAD.....................................................................................................................4-36 4.4.2 RVAD.....................................................................................................................4-37 4.4.3 Bi-LVAD................................................................................................................4-38 4.4.4 Bi-VAD ..................................................................................................................4-39

4.5 Discussion ................................................................................................. 4-40 4.6 Conclusion ................................................................................................ 4-41

Chapter 5 VAD Experimental Evaluation ........................................................... 5-1

5.1 Introduction................................................................................................. 5-1 5.2 Experimental Method.................................................................................. 5-3

5.2.1 Prototype Pump Construction...................................................................................5-4 5.2.2 Equipment / Instrumentation / Software / Code .......................................................5-5 5.2.3 Experimental Design ................................................................................................5-7 5.2.4 Experimental Procedure ...........................................................................................5-9 5.2.5 Calibration ..............................................................................................................5-11 5.2.6 Experimental Configurations..................................................................................5-13

5.3 Results ....................................................................................................... 5-15 5.3.1 Non-Pulsatile ..........................................................................................................5-15 5.3.2 Pulsatile ..................................................................................................................5-26

5.4 Discussion ................................................................................................. 5-26 5.4.1 Experimental Method .............................................................................................5-26 5.4.2 Non-Pulsatile Operation .........................................................................................5-26 5.4.3 Pulsatile Operation .................................................................................................5-26 5.4.4 Application to Bearing System Design...................................................................5-26

5.5 Conclusion ................................................................................................ 5-26 5.5.1 Force and Performance...........................................................................................5-26 5.5.2 Bearing Design .......................................................................................................5-26

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Chapter 6 VAD Design Summary....................................................................... 6-26 6.1 Introduction ............................................................................................... 6-26 6.2 Bi-LVAD................................................................................................... 6-26

Overview ................................................................................................................................ 6-26 6.3 Bi-VAD ..................................................................................................... 6-26

Overview ................................................................................................................................ 6-26 Chapter 7 Conclusions and Future Research .................................................... 7-26

7.1 Conclusions ............................................................................................... 7-26

7.1.1 Literature Review ................................................................................................... 7-26 7.1.2 Mock Circulation Loop .......................................................................................... 7-26 7.1.3 Centrifugal VAD Design and Development........................................................... 7-26 7.1.4 VAD Experimental Evaluation............................................................................... 7-26 7.1.5 VAD Design Detail ................................................................................................ 7-26

7.2 Future Research......................................................................................... 7-26 7.2.1 Mock Circulation Loop .......................................................................................... 7-26 7.2.2 Centrifugal VAD Development.............................................................................. 7-26

Appendices

Appendix A - A Survey of Current Mechanical Assist Devices

Appendix B - Centrifugal VAD Design

Appendix C - Impeller Hydraulic Force Calculation

Appendix D - SIMULINK and MATLAB Code

Appendix E - Calibrations

Appendix F - Magnetic Bearing Investigation

Appendix G - VAD Design Detail

References

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List of Figures Figure 2-1 – Inlet (Left) and Outlet (Right) Devices......................................................................... 2-27 Figure 2-2 –Double Volute Configuration ........................................................................................ 2-28 Figure 2-3 - Non-Dimensional Performance of Rotary Blood Pumps ............................................. 2-29 Figure 2-4 – Experimental Measurement of Radial Thrust Magnitude and Direction ...................... 2-35 Figure 2-5 – Values of Radial Force Factor (K) for Eccentrically Located Impellers in a Concentric Casing ............................................................................................................................................... 2-35 Figure 2-6 – Monopivot Magnetic Suspension Pump....................................................................... 2-38 Figure 2-7 – Radial Restoration Using Permanent Magnets ............................................................. 2-38 Figure 2-8 – Radial Magnetic Motor-Bearing................................................................................... 2-39 Figure 2-9 – Principle of Passive Restoring Force............................................................................ 2-40 Figure 2-10 – Axial Magnetic Motor-Bearing .................................................................................. 2-41 Figure 3-1 – Circulatory System ......................................................................................................... 3-3 Figure 3-2 – Circulatory Pressure Distribution ................................................................................... 3-3 Figure 3-3 – Phases and Pressure for one Cardiac Cycle ................................................................... 3-4 Figure 3-4 – Pressure Distribution within the Chambers of the Heart (Rest)...................................... 3-8 Figure 3-5 – Left Heart Pressure Distribution..................................................................................... 3-8 Figure 3-6 – Systemic Pressure Distribution..................................................................................... 3-10 Figure 3-7 – Pulmonic Pressure Distribution .................................................................................... 3-10 Figure 3-8 – Compliance (systole and diastole) ................................................................................ 3-11 Figure 3-9 – Definition of Compliance ............................................................................................. 3-11 Figure 3-10 – Development of Ventricular Stroke Volume .............................................................. 3-16 Figure 3-11 – Effect of Volume, Vascular Compliance and Resistance on PRA ............................... 3-20 Figure 3-12 – Cardiac Function Curves ............................................................................................ 3-21 Figure 3-13 – Interaction between Cardiac and System Function Curves ........................................ 3-21 Figure 3-14 – Compensatory mechanisms in response to heart failure........................................... 3-21 Figure 3-15 – Basic Test Loop.......................................................................................................... 3-25 Figure 3-16 – Advanced Mock Circulation Loop ............................................................................. 3-26 Figure 3-17 – (Pantalos, Koenig et al. 2004) Mock Circulation Loop .............................................. 3-27 Figure 3-18 –(Wu, Allaire et al. 2004) Mock Circulation Loop ....................................................... 3-27 Figure 3-19 – ASTM F1841Test Loop (ASTM_F1841-97 1998).................................................... 3-33 Figure 3-20 – SIMULINK Model for Complete Circulation ............................................................ 3-35 Figure 3-21 – Resting Pressure Distribution ..................................................................................... 3-36 Figure 3-22 – Left Heart Failure Pressure Distribution................................................................... 3-36 Figure 3-23 – Resting Perfusion........................................................................................................ 3-36 Figure 3-24 – Left Heart Failure perfusion ....................................................................................... 3-36 Figure 3-25 – Mock Circulation Rig ................................................................................................. 3-38 Figure 3-26 – Top View Mock Circulation Loop with listed parameters ......................................... 3-40 Figure 3-27 – Structural components of the simulated beating heart ................................................ 3-41

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Figure 3-28 – Control and monitor of cardiac function .................................................................... 3-42 Figure 3-29 – Compliance Chamber ................................................................................................. 3-45 Figure 3-30 – Proportional Pinch Valve ........................................................................................... 3-45 Figure 3-31 – Electromagnetic Flowmeter........................................................................................ 3-48 Figure 3-32 – Experimental Technique............................................................................................. 3-50 Figure 3-33 – Pressure distribution throughout the vascular tree for systemic (a) and pulmonic (b) circulation` systems for resting conditions........................................................................................ 3-54 Figure 3-34 – Pressure distribution throughout the vascular tree for systemic (a) and pulmonic (b) circulation systems for heart failure conditions. ............................................................................... 3-54 Figure 3-35 – Perfusion Rate for Rest............................................................................................... 3-54 Figure 3-36 – Perfusion Rate for Left Heart Failure. ....................................................................... 3-54 Figure 3-37 – Systemic Pressure Distribution (Exercise) ................................................................. 3-56 Figure 3-38 - Perfusion Rate for systemic circulation (Exercise) ..................................................... 3-56 Figure 4-1 – Volute velocity assuming constant momentum (a) and constant velocity (b) ................ 4-8 Figure 4-2 – –Volute Configurations ................................................................................................. 4-9 Figure 4-3 – Types of volute cross sections ...................................................................................... 4-10 Figure 4-4 – Volute design parameters (Lazarkiewicz and Troskolanski 1965)............................... 4-11 Figure 4-5 – Taper Angle Vs Throat Velocity .................................................................................. 4-11 Figure 4-6 - Non-Dimensional Performance for Different Discharge Blade Angles ....................... 4-12 Figure 4-7 – Variations in Static Radial Thrust of Common Volute Types...................................... 4-16 Figure 4-8 – Single Volute Fluid Velocity at Various Flow Conditions (Q). ................................... 4-17 Figure 4-9 - Radial Thrust Direction at Operation Below Design Capacity ..................................... 4-18 Figure 4-10 - Radial Thrust Direction at Operation Above Design Capacity ................................... 4-19 Figure 4-11 – Flow Separation Regions in a Double Volute at Part Capacities................................ 4-20 Figure 4-12 – Modified Concentric Volute....................................................................................... 4-21 Figure 4-13 - Radial Thrust Direction for Single Volutes................................................................. 4-23 Figure 4-14 – Variations in Static Axial Thrust in all Volute Types ................................................ 4-25 Figure 4-15 – Variations in Static Axial Thrust in an Closed Type Impeller.................................... 4-26 Figure 4-16 – Variations in Static Axial Thrust in a Semi-Open Type Impeller............................... 4-26 Figure 4-17 – Techniques for Balancing Axial Thrust...................................................................... 4-27 Figure 4-18 – LVAD Pump .............................................................................................................. 4-37 Figure 4-19 – RVAD Pump .............................................................................................................. 4-37 Figure 4-20 – Bi-LVAD Pump ......................................................................................................... 4-38 Figure 4-21 – Bi-VAD Pump............................................................................................................ 4-39 Figure 5-1 – Experimental Process Diagram ...................................................................................... 5-2 Figure 5-2 –Force Test Rig ................................................................................................................. 5-8 Figure 5-3 – Pressure Tapping Locations ........................................................................................... 5-9 Figure 5-4 – Experimental VAD configurations............................................................................... 5-12 Figure 5-5 – LVAD Family of Pump Performance and Efficiency Curves ..................................... 5-14 Figure 5-6 – LVAD Performance with Constant Speed and Axial Clearance Variation .................. 5-16 Figure 5-7 – LVAD Non-Dimensional Pump Performance Curves.................................................. 5-17

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Figure 5-8 – LVAD Non-Dimensional Power Curves ...................................................................... 5-18 Figure 5-9 – RVAD Family of Pump Performance Curves .............................................................. 5-19 Figure 5-10 – RVAD Non-dimensional Pump Performance Curves ................................................ 5-19 Figure 5-11 – Radial Thrust in single (a), double (b) and circular (c) volutes. ................................. 5-21 Figure 5-12 – Radial Thrust Direction & Magnitude Integrated from Single Volute Pressures Readings at Shut-off, Design and Maximum Flow Capacities. ........................................................................ 5-22 Figure 5-13 – Axial Thrust on Impeller Types.................................................................................. 5-23 Figure 5-14 – Comparison of Axial Thrust in a Semi-Open single and double impeller .................. 5-23 Figure 5-15 – Reaction Torque in Single Volute (22.5).................................................................... 5-24 Figure 5-16 – Comparison of measured Radial thrust with theoretical prediction............................ 5-25 Figure 5-17 – Comparison of measured axial thrust with theoretical prediction .............................. 5-26 Figure 5-18 – Direction of Radial Thrust in Single, Double and Circular Volutes at 2400rpm........ 5-26 Figure 5-19 – Radial thrust direction of comparative specific speeds in a single volute .................. 5-26 Figure 5-20 – Untreated heart failure pressure for systemic (i) and pulmonary (ii) circulation. ....... 5-26 Figure 5-21 – LVAD supported pressure for systemic (i) and pulmonary (ii) circulation. .............. 5-26 Figure 5-22 – Transition of perfusion from Heart Failure to LVAD Support ................................... 5-26 Figure 5-23 – Non-Pulsatile LVAD operation in the pulsatile LHF environment ............................ 5-26 Figure 5-24 – Non-Pulsatile LVAD efficiency in the pulsatile LHF environment ........................... 5-26 Figure 5-25 – Fluctuation of force with systole and diastole ............................................................ 5-26 Figure 5-26 –Forces from pulsatile and non-pulsatile test results ..................................................... 5-26 Figure 5-27 – Systemic (a) and Pulmonic (b) BI-HF pressure distribution....................................... 5-26 Figure 5-28 – Systemic (a) and Pulmonic (b) BI-VAD supported pressure distribution ................. 5-26 Figure 5-29 – Perfusion Rate for Rest............................................................................................... 5-26 Figure 5-30 – Perfusion Rate for Left Heart Failure. ....................................................................... 5-26 Figure 5-31 – Pulsatile Hydraulic LVAD Performance in Various Heart Function Situation .......... 5-26 Figure 6-1 – Exploded View of Bi-LVAD Design............................................................................ 6-26 Figure 6-2 – Schematic Cross Sectional View of the Bi-LVAD design .......................................... 6-26 Figure 6-3 – Perspective View of the Bi-LVAD Design................................................................... 6-26 Figure 6-4 – Exploded View of Bi-VAD Design.............................................................................. 6-26

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List of Tables Table 2-1 - The index of hæmolysis of blood pumps.......................................................................... 2-8 Table 2-2 – Current Mechanical Assist Devices ............................................................................... 2-18 Table 2-3 – Dimensionless Coefficients ........................................................................................... 2-29 Table 2-4 – Comparison of Self Bearing Motor Operations Characteristics..................................... 2-43 Table 3-1 – Blood Volume Distribution ............................................................................................. 3-3 Table 3-2 – Resting Pressure Distribution throughout the Vascular Tree......................................... 3-10 Table 3-3 –Mean Circulatory Pressures for Rest/LHF/Exercise ....................................................... 3-10 Table 3-4 – Arterial Compliance....................................................................................................... 3-12 Table 3-5 – Venous Compliance....................................................................................................... 3-13 Table 3-6 –Systemic Vascular Resistance ........................................................................................ 3-14 Table 3-7 –Pulmonic Vascular Resistance........................................................................................ 3-15 Table 3-8 – Parameter affecting cardiac output ................................................................................ 3-17 Table 3-9 – Properties of Pumping Mediums ................................................................................... 3-32 Table 3-10 – Valve Settings for Various Circulation Configurations ............................................... 3-41 Table 3-11 – Chamber characteristics for Resting Compliance Values ............................................ 3-45 Table 3-12 – Mock Circulation Rig Component Inherent Resistance Values................................... 3-45 Table 3-13 – Physiological flow conditions...................................................................................... 3-48 Table 3-14 – Mock Circulation Loop Parameter............................................................................... 3-52 Table 3-15 – Comparison of Mock Circulation Loop Pressure Results............................................ 3-60 Table 3-16 – Systemic (SVR) and Pulmonary (PVR) Resistance Comparison ................................ 3-63 Table 3-17 – Systemic (SAC) and Pulmonary (PAC) Compliance Comparison .............................. 3-63 Table 3-18 – Comparison of Mock Circulation Loop Perfusion Results .......................................... 3-65 Table 4-1 – Fluid Properties.............................................................................................................. 4-33 Table 4-2 – Design Output Pressure ................................................................................................. 4-33 Table 4-3 - Design Flow Rates.......................................................................................................... 4-33 Table 4-4 – Reynolds Number (Re) .................................................................................................. 4-34 Table 4-5 – Specific Speed (Ns) ........................................................................................................ 4-35 Table 4-6 – Specific Capacity (qs) .................................................................................................... 4-35 Table 4-7 – Specific Head (hs) .......................................................................................................... 4-35 Table 4-8 – LVAD Impeller and Volute Configurations .................................................................. 4-36 Table 4-9 – RVAD Impeller and Volute Configurations .................................................................. 4-37 Table 4-10 – Bi-LVAD Impeller and Volute Configurations ........................................................... 4-38 Table 4-11 – Bi-VAD Impeller and Volute Configurations.............................................................. 4-39 Table 5-1 – LVAD Impeller/Volute Experimental Configurations................................................... 5-13 Table 5-2 – RVAD Impeller/Volute Experimental Configurations .................................................. 5-13 Table 5-3 – BVAD Impeller/Volute Experimental Configurations .................................................. 5-13 Table 5-4 - Design Point Rotational Speed (RPM) of LVAD Impeller/Volute Configurations ....... 5-15 Table 5-5 - Design Point Rotational Speed (RPM) of RVAD Impeller/Volute Configurations ....... 5-20

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Table 5-6 – Radial trust at design point (5L/min @ 100mmHg)....................................................... 5-22 Table 5-7 – Axial thrust at design point ............................................................................................ 5-23 Table 5-8 – Comparison of Heart Failure to LVAD Support Component Pressures ........................ 5-26 Table 5-9 – Pressure and flow conditions for diastolic and systolic periods..................................... 5-26 Table 5-10 – LVAD Support Force/Torque Magnitudes at Systole and Diastole............................. 5-26

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Nomenclature AMB Active Magnetic Bearing LVP Left Ventricular Pressure A1,Awr Impeller Inner Diameter Area LVPED

Left Ventricular End Diastolic Pressure

A2, A3 Impeller Outer Diameter Area MAP Mean Arterial Pressure AB Area exposed back shroud pressure MPAP Mean Pulmonary Arterial Pressure Ae Impeller Eye Area MSCBP Magnetically Suspended Centrifugal

Blood Pump AF Front Shroud Ns Specific Speed As,Ash, Ah

Shaft Area P Pressure (Head)

AoC Aortic Compliance PAC Pulmonary Arterial Compliance AoP Aortic Pressure PAP Pulmonary Arterial Pressure Ath Throat Area PP Pulse Pressure B Magnetic Field Strength PRU Pulmonary Resistance Units BEP Best Efficiency Point PVC Pulmonary Venous Compliance Bpm Beats per Minute PVP Pulmonary Venous Pressure b1 Inlet Blade Height PVR Pulmonary Vascular Resistance b2 Outlet Blade Height P1 Pressure at Inner Diameter B2 Impeller Width including shrouds PD Head, Discharge Pressure C Compliance Pdias Diastolic Pressure Ca Arterial Compliance Pmc Mean Circulatory Pressure CI Cardiac Index Ps , ps Suction Pressure CO Cardiac Output Psys Systolic Pressure CPB Cardio-pulmonary Bypass Q Flow Rate (Capacity) CVP Central Venous Pressure Qn Normal Capacity (B.E.P) c1 Fluid Inlet/Eye Velocity qs Specific Capacity cm2 Meridional Outlet Velocity R Resistance Co Fluid Inlet/Eye Velocity RAP Right Atrial Pressure cu1 Inlet Tangential Velocity RVP Right Venous Pressure cu2 Outlet Tangential Velocity SV Stroke Volume cu3 Throat Velocity SVC Systemic Venous Compliance Cv Venous Compliance SVP Systemic Venous Pressure D1 Impeller Inlet Diameter SVR Systemic Vascular Resistance D2 Impeller Outlet Diameter T Net/Resultant Axial Thrust D3 Base Circle Diameter Ta Net/Resultant Axial Thrust EDV End Diastolic Volume TPR Total Peripheral Resistance ESV End Systolic Volume u2 Outer Diameter Peripheral Velocity F Momentum change force ush, us Shaft Peripheral Velocity FT Net/Resultant Axial Thrust uwr Inner Diameter Peripheral Velocity Fr Radial Force (lbf) VR Venous Return H Head, Discharge Pressure HR Heart Rate Hp Head, Discharge Pressure Hv Pressure at Inner Diameter hs Specific Head α2’ Actual Fluid Discharge Angle I Current in Coil αv Volute Spiral Angle K Radial Thrust Factor β1 Impeller Inlet Angle KAvb Avg % pressure on Back Shroud β2 Impeller Discharge Angle KAvf Avg % pressure on Front Shroud ∆p Developed Impeller Pressure Kth Radial Thrust Factor δ Diffuser Angle L Effective length of wire γ , sp.gr Specific Gravity (1.0 for water) LAP Left Atrial Pressure LHF Left Heart Failure

xv

Statement of Authorship The work contained in this thesis has not been previously submitted for a degree or

diploma at any other higher education institution. To the best of my knowledge and

belief, the thesis contains no material previously published or written by any other

person except where due reference is made.

Daniel Timms

2005

xvi

Acknowledgements I would first like to acknowledge the unrelenting and invaluable support of my

parents. The completion of this research is a testament to their support throughout

my entire candidature, from brainstorming sessions with my father to the use of our

growing home workshop. The significance of this field was highlighted after the

repair and eventual replacement of his mitral valve in 2004 by the Transplant

Services Department of the Prince Charles Hospital, the team with which research

collaboration was achieved at the commencement of the project.

I cannot describe the extent to which my gratitude is extended to the Mechanical

Engineering Electronics Technician, Mark Hayne. His job description does not begin

to pay justice to the range of contributions made to assist in my successful

completion. Countless hours were spent milling over the workshop whiteboard,

which were crucial to the identification of research avenues to pursue. His knack at

pointing out the bright side was an appreciated characteristic essential for my ability

to face, and then overcome adversities.

To my supervisor, Andy Tan, for his enthusiasm in an exciting field of research. His

supervision ultimately led to the unconstrained experimentation required for the

development of new designs.

To Mark Pearcy, the ever present leader of Medical Engineering Research at QUT,

who possesses the ability to critically apply engineering principles to all fields of

research. My future career will be no doubt influenced by the goal of trying to

experience success to a similar magnitude.

Appreciation is also extended to Keith McNeil, Andrew Galbraith and colleagues at

the Prince Charles Hospital for their consistent and unrelenting enthusiasm about the

project as a whole. The supply of research funding was critical in the development

of prototype pumps and mock circulation testing facility.

xvii

I have been fortunate enough to have had the support of close friends and assistance

from colleagues throughout my candidature. I wish to extend sincere thanks to the

following people:

• Lydia Balzat for her inspiring attitude and support throughout my

candidature,

• Katie McNickle for the work ethic developed in undergrad that led to the

opportunity to tackle postgraduate studies,

• Nicky McAllister, for a part-time job that provided financial security as well

as an environment suitable for countless hours of extra research,

• Andrew Cliffe for his assistance in the flow visualisation experimentation,

• Aaron Cahill, Ben Breise, Noboyuki Kurita and Takahito Tokomoto for their

work and valuable assistance with the various magnetic bearing

configurations,

• Professor Yohji Okada, for the opportunity to study Magnetic Bearings in his

laboratory for 3 months in 2001,

• Peter Ridley for his experience in control used in the development of the

Mock Loop simulation,

• Wim Dekkers for his expertise in fluid dynamics.

My initial interests in medical devices stemmed from my undergraduate studies in

medical engineering. Although my degree was completed in Mechanical

Engineering, there was little doubt I would once again ply my learned engineering

skills to the pursuit of improving quality of life. For this, I must again thank my

supervisor, Andy Tan, for the opportunity to pursue cardiac assist devices at QUT.

The team of researchers developed, and funding attracted since the projects inception

is a tribute to each individual mentioned for assisting and contributing to the ultimate

goal of a viable ‘artificial heart’.

Chapter 1 - Introduction

1-1

Chapter 1

Introduction

Heart disease is the developed world’s biggest killer, with over 800,000 new cases

reported each year. Unfortunately, less than 3000 heart transplant procedures are

performed annually due to a shortage of donor hearts (Rose, Gelijns et al. 2001). This

predicament has accelerated the development of mechanical circulatory assistance as

an alternative treatment.

Studies have revealed that patients receiving mechanical devices to assist a failing

left ventricle have extended lives and improved quality of life compared to recipients

of optimal drug therapy. However, these patients succumb to various ailments

associated with currently approved pulsatile type devices inside two years (Rose,

Gelijns et al. 2001). Therefore, current research is focused on third generation

continuous flow rotary pumps that employ magnetic or hydrodynamic techniques to

suspend the impeller within the pump casing (Baloh, Allaire et al. 1999; Watterson,

Woodard et al. 2000). These systems reduce blood damage, component wear and

potentially increase device lifetime beyond 10 years.

This thesis outlines the design and progressive development of two novel, third

generation, continuous flow, double sided centrifugal impeller pumps; a bi-left

ventricular assist device (Bi-LVAD) and single bi-ventricular assist device (Bi-

VAD). The Bi-LVAD is designed to assist the function of the left ventricle only,

while the single Bi-VAD simultaneously augments the function of both left and right

ventricles of patients suffering bi-ventricular heart failure. Alternative techniques for

the latter condition require the implantation of two devices in such patients (Nose

and Furukawa 2004), increasing overall system size and reducing the access of

women and children to this vital technology.


1-2

Magnetic and hydrodynamic bearing techniques were identified to suspend the

double sided impeller in each miniature centrifugal type device, thus overcoming

conventional pump limitations of friction and wear at the drive shaft, seal and

bearing interface. This will greatly increase the durability of the pump and reduce

both heat generation and red blood cell damage. Complete impeller suspension

requires the generation of bearing forces to overcome the dynamic and static

disturbance forces experienced during operation. By designing the bearing with

sufficient load capacity, the incidence of impeller touchdown can be minimised.

Furthermore, the reduction of static load will reduce the power requirements of the

magnetic bearing system to maintain the impeller in a central position.

Although the final designs are still subject to iterative refinement, this thesis

describes the processes undertaken to develop each device to their current stage.

Progressive experimentation was conducted and conclusions drawn from a number

of medical and engineering fields, such as flow visualization, hemodynamics and

heart failure, rotary pump design and testing, and magnetic motor bearing design.

The requirement for in-vitro device evaluation prompted the design and construction

of a new systemic and pulmonary mock circulation loop, including pulsatile left and

right ventricles coupled with vascular compliances and resistances. Since most VAD

centers focus on left ventricular assistance, no mock circulation loop existed that

could test the function of left and/or right assistance. The custom developed loop

was therefore designed to replicate the hemodynamic conditions of the entire

cardiovascular system in response to normal heart function and varying levels of left

and/or right heart failure. This was achieved by exclusively controlling

hemodynamic parameters such as heart function (rate, contractility) and vascular

tone (resistance, compliance), a feature not exclusively afforded in an in-vivo setting.

This small and compact rig provides the ability to evaluate the hemodynamic effect

of Left-, Right and Bi- ventricular assist device in-vitro. It has the potential to reduce

device evaluation costs by simulating the natural circulatory system, and thus

provide valuable cardiovascular device performance feedback prior to monetary and

time expensive and in-vivo animal trials.


1-3

A centrifugal pump design procedure was then formulated to design the components

of each VAD. A number of single sided impeller/volute configurations were

constructed as individual left and right components of the Bi-VAD design, to

produce the required pressure and flow performance to assist the systemic and

pulmonary circulatory systems.

The performance of a rotary blood pump is gauged by the success of the device to re-

establish cardiovascular hemodynamic parameters (pressure and flow) from

pathological to normal levels. This capability is determined by the pump

performance characteristic curves. The gradient of the curve influences the response

of the device to changing physiological conditions (exercise, rest, etc).

Device operation results in a fluid pressure distribution within the pump casing,

which if unbalanced, lead to axial and radial hydraulic thrust acting on the impeller.

A unique method for measuring the magnitude and direction of impeller static

hydraulic thrust during device operation was therefore developed. Force

characteristics are valuable for impeller suspension design of third generation

magnetic or hydrodynamic bearing devices. Specifically, the force magnitude

determines the required magnetic or hydrodynamic forces to keep the impeller

central, and minimise contact with the casing.

Each of the constructed LVAD and RVAD components were inserted into the mock

circulation loop configured for non-pulsatile operation, to record performance

characteristic curves. An LVAD configuration was then inserted into the loop

configured for left heart failure to assess its ability to reduce pulmonary congestion,

and restore perfusion levels from pathological to normal levels. Impeller force

characteristics were simultaneously recorded during all tests involving each LVAD

configuration. The influence of failing heart pulsatility on the magnitude and

direction of these hydraulic forces was determined


1-4

The results presented in this thesis are used to recommend the appropriate

impeller/volute configurations for designs incorporating magnetic, hydrodynamic

and mechanical type bearings to increase device lifetime. The force information was

then specifically used as criteria for magnetic motor-bearing investigations, to

determine a suitable configuration for incorporation in each of the proposed Bi-

LVAD and Bi-VAD designs.

Finally, the research conducted and conclusions drawn in the thesis were used to

design the current Bi-LVAD and Bi-VAD concepts. The proposed Bi-LVAD

incorporates symmetrical blade profiles on each side of the magnetically suspended

double impeller. The device assists the function of the left ventricle only with

symmetrical axial pressure distribution and elimination of stagnant regions beneath

the impeller. These features improve axial touchdown capacity and reduce thrombus

formation respectively. The proposed Bi-VAD incorporates different blade designs

on each side of the magnetically and hydrodynamically suspended double impeller,

to augment the function of both the left and right cardiac chambers.


1-5

1.1 Significance

The significance of this research relates to the development of two novel centrifugal

heart assist devices. These miniature devices can afford smaller patients access to

VAD technology, potentially saving many lives each year.

Successful development of these innovative VAD’s will provide an alternative to

heart transplantation plantation for many cardiac patients worldwide. No longer

would these patients need to wait for the untimely death of a donor to provide a

suitable heart. Instead, these latest generation devices would be available

immediately, and be almost universally compatible. They have the potential to

dramatically increase a patient’s expected lifetime, and to deliver them a higher

quality of life. This translates to a renewed ability to leave the confines of the

hospital and return to daily activities. Furthermore, this situation can potentially ease

the strain on worldwide health care systems caused by cardiovascular disease, as

hospitalisation costs may be reduced, or reallocated to other critical care. Patients can

be transferred from expensive Intensive Care Units to Wards, and ultimately to the

comfort of their own home. VAD therapy can also reduce the use of expensive and

continuous pharmaceuticals, in the form of immunosuppressant drugs associated

with open heart transplantation, and BETA Blockers/ACE inhibitors related to heart

failure treatment.

The development of a mock circulation loop that has the capability of reproducing

the hemodynamics of a patient experiencing a range of cardiovascular diseases will

provide valuable insights into each cardiovascular device performance prior to in-

vivo trials and human implantation.

The identification and characterisation of static hydraulic force on a VAD impeller

operating in a pulsatile environment will help to determine the required magnetic or

hydrodynamic bearing load capacity to minimise impeller touchdown. By

investigating numerous impeller and volute configurations, the most appropriate

combination may be selected to suit each bearing design.


1-6

1.2 Objectives

The main objective of this thesis was to describe the design, development and

evaluation process of two novel rotary type ventricular assist devices, a Left

Ventricular Assist Device (LVAD) and Bi-Ventricular Assist Device (Bi-VAD).

The development of a pulsatile mock cardiovascular loop was specifically required to

achieve the secondary objective of static hydraulic force measurement on the

impeller of a centrifugal blood pump. A suitable test facility was then required to

isolate impeller forces in the pulsatile operating environment, to determine the

required bearing load capacity to minimise impeller touchdown and reduce magnetic

bearing power.

Substantial clinical input was merged with engineering principles for the purpose of

creating this cardiovascular test rig to reproduce the pulsatile human cardiovascular

system. This mock circulation loop was required to investigate clinical circulatory

hemodynamics in response to a variety of acute or chronic heart failure conditions,

by recreating their symptomatic effect on hemodynamic parameters (vascular

resistance, contractility, heart rate). Ultimately, the objective of this rig was to

facilitate VAD testing prior to in-vivo animal and clinical trials, by assessing the

device’s ability to return the circulatory hemodynamics to normal conditions.


1-7

1.3 Aims

The specific aims of this thesis are described with reference to each chapter:

Literature Review

1. The review of the literature aimed to detail the requirements for VAD design

and support the pursuit of Bi-Ventricular assistance.

2. Research into active magnetic motor-bearings aimed to provide characteristic

information about each configuration type (radial and axial), such as max

load capacity in the supported DOF, efficiency and rotational speed.

3. The flow visualisation background research attempted to highlight regions for

potential blood trauma in centrifugal blood pumps.

Mock Circulation Loop

1. To design and construct a systemic and pulmonary mock circulation loop that

could recreate the hemodynamic conditions of a normal and failing heart at

rest.

2. To create a computer simulation model to validate and assist in the design,

construction and validation of the invitro circulation loop.

Centrifugal VAD Design and Development

1. To produce a customized method for VAD design iterations

2. To create a number of impeller and volute configurations for evaluation.

VAD Experimental Evaluation

1. To identify the pump performance and force characteristics of each

impeller/volute configuration in non-pulsatile and pulsatile mock circulation

loops.

2. To select the best pump impeller/volute configurations to both improve

performance within the cardiovascular environment, and magnetic bearing

load capacity to minimise impeller touchdown.

VAD Design Detail

1. To provide a detailed discussion of each VAD design using the results

obtained and conclusions drawn from the preceding chapters.


1-8

1.4 Thesis Outline

This thesis is presented in six chapters describing the design and progressive

development of two ventricular assist devices with the aid of a suitable mock

circulation loop to evaluate their performance in an in-vitro pulsatile heart failure

environment.

Chapter 2 continues with a review of the literature and background research into this

emerging field, allowing the reader to understand and appreciate the importance of

cardiac assist device development, and the niche that the proposed designs occupy.

Within this review, the design criteria for each VAD, from potential blood damaging

effects to the hemodynamic requirements of the device, are discussed.

A summary of two background research topics are also included in this chapter. The

findings from a study on magnetic motor bearing technology outline the possible

suspension configurations for incorporation in both LVAD and Bi-VAD

applications, while the flow visualisation information attempts to highlight the

general regions of potential blood trauma in centrifugal blood pumps.

A suitable testing facility was required to evaluate the performance of each VAD

design configuration. A complete mock circulation loop, including systemic and

pulmonary circuits in series, was constructed and detailed in Chapter 3. This rig

incorporated passive filling atria and pneumatically actuated ventricles with the

ability to alter function in order to simulate the hemodynamic conditions of varying

degrees of left and/or right heart failure. A first order SIMULINK computer model of

the circulation system was also simultaneously developed to assist in rig construction

as well as result verification.

Chapter 4 then presents the components required for centrifugal pump design. A

variety of methods were amalgamated in Appendix B to produce a custom procedure

used to design the first and future VAD impeller and volute iterations . A number of

volute and impeller configurations were designed with the intention of evaluating

each and selecting the most suitable for inclusion in each final VAD design. The Bi-

VAD was initially designed as two single LVAD and RVAD pumps, under the

premise that results obtained from these configurations will assist in the final

iteration design. Details of these individual designs conclude the chapter.


1-9

With the proposed method of impeller suspension relying on magnetic force

generation to levitate the impeller within the pump cavity, the magnitude of these

force requirements should be ascertained in order to complete the bearing design.

Knowledge of the expected magnitudes in each degree of freedom determines the

required bearing load stiffness to minimise impeller touchdown.

The assessment of impeller force and performance characteristics experienced by

each impeller/volute configuration is conducted in Chapter 5. Impellers of varying

discharge angle were supported within the single, double and circular volute casings

by a conventional shaft, seal and bearing system assembly, and driven by a small

electric motor.

Performance characteristics were measured initially in a non-pulsatile mock

circulation environment, before the hemodynamic effect of VAD insertion in the

complete pulsatile mock circulation rig experiencing left heart failure, such as the

relief of pulmonary congestion and return of sufficient perfusion was assessed.

Particular focus was placed on the impeller discharge angle and axial actuation of the

semi-open type impeller

Simultaneous measurement of radial (x,y) and axial (z) hydraulic forces imposed on

the impeller by differential hydraulic pressures within the volute casing was also

undertaken. The results were then matched to force calculation equations to

theoretically predict axial and radial thrust for VAD designs. This information is

crucial for the impeller suspension design to minimise impeller touchdown.

Consequent recommendations are made for suitable impeller/volute configurations

for magnetic, hydrodynamic and mechanical bearing suspension.

Chapter 6 summarises the current VAD designs, taking into consideration

conclusions drawn from the research presented in the thesis. Functionality is

discussed while advantages and disadvantages are critically provided to assess the

capability of each VAD to accomplish the task of ventricular assistance.

The thesis is concluded in Chapter 7 by addressing the thesis aims, and summarising

and reiterating the chapter contributions to each specific VAD design.


1-10

1.5 Thesis Outcomes and Contributions

The work presented in this thesis describes a significant development in the design

and in-vitro testing of ventricular assist devices, and encompasses new and

innovative concepts that have led to:

1. Two novel ventricular assist device design concepts that include;

a. A Bi-LVAD device incorporating a double suction centrifugal

impeller.

b. The combination of a rotary LVAD and RVAD into a single impeller

Bi-VAD device.

2. The development of a model simulation and physical complete pulsatile

systemic and pulmonary mock circulation loop which;

a. Can simulate the effect of varying levels of heart disease on the

systemic and pulmonary circulatory system.

b. Is suitable for evaluating the hemodynamic performance of

cardiovascular devices.

3. A design procedure to develop centrifugal type ventricular assist devices

which

a. is an amalgamation of various pump design recommended procedures

b. provides a step by step outline for the design of each component

4. A unique method for measuring hydraulic forces in impeller/volute

configurations, with the results used to;

a. Determine the load requirements of magnetic or hydrodynamic

bearings to minimise impeller touchdown in a pulsatile environment.

b. Suggest the most appropriate bearing configuration for each design.

c. Select the most appropriate empirical formula for hydraulic force

prediction.


1-11

1.6 Publications

The full text of each publication listed below is located on the accompanying CD-

ROM.

1.6.1 Peer Reviewed Journal Publications

Timms, D.L., M. Hayne, A. Galbraith, and K. McNeil, A Complete Mock

Circulation Loop for the Evaluation of Left- Right- and Bi- Ventricular Assist

Devices. Artificial Organs, 2005. 29(7): p. 564–571.

Timms, D.L., M. Hayne, A.C.C. Tan, and M.J. Pearcy, LVAD Pump Performance

And Force Characteristics In A Pulsatile Complete Mock Circulation Loop. Artificial

Organs, 2005. 29(7): p. 572-580

Timms, D.L., A.C.C. Tan, M.J. Pearcy, K. McNeil, and A. Galbraith, Hydraulic

Force and Impeller Evaluation of a Centrifugal Heart Pump. Journal of the Korean

Society of Marine Engineers, 2004. 28(2): p. 376-381.

Tan, A.C.C., D.L. Timms, M.J. Pearcy, K. McNeil, and A. Galbraith, Experimental

Flow Visualisation of an Artificial Heart Pump. Journal of the Korean Society of

Marine Engineers, 2004. 28(2): p. 210-216.

1.6.2 Published Conference Proceedings

Tokumoto, T., D.L. Timms, H. Kanebako, K. Matsuda, and Y. Okada. Development

of Lorentz Type Self Bearing Motor. in 8th International Symposium on Magnetic

Bearings. 2002. Mito, Japan.

Timms, D.L. and A.C.C. Tan. Flow Visualisation of a Centrifugal Pump: The

Relationship to an Artificial Heart. in Australasian Association for Engineering

Education. 2002. Canberra, Australia.

Timms, D.L. and A.C.C. Tan. Hydraulic Force and Impeller Evaluation of a

Centrifugal Heart Pump. in International Symposium on Combustion Engine and

Marine Engineering (ISCEM 2003). 2003. Busan, Korea.

Tan, A.C.C. and D.L. Timms. Experimental Flow Visualisation of an Artificial

Heart Pump. in International Symposium on Combustion Engine and Marine

Engineering (ISCEM 2003). 2003. Busan, Korea.


1-12

1.6.3 Published Abstracts

Timms, D. L. Complete Mock Circulation Loop For The Evaluation Of

Cardiovascular Devices in Heart Failure and Rotary Blood Pump Summit, 2004.

Cleveland, Ohio, USA.

Timms, D. L. LVAD Pump Performance and Force Characteristics in a Pulsatile

Mock Circulation Loop in Heart Failure and Rotary Blood Pump Summit, 2004.

Cleveland, Ohio, USA.

Timms, D.L., A.C.C. Tan, and M.J. Pearcy. Flow Visualisation of a Centrifugal

Blood Pump. in World Congress on Medical Physics and Biomedical Engineering

(WC2003). 2003. Sydney, Australia.

Timms, D.L., A.C.C. Tan, M.J. Pearcy, K. McNeil, and A. Galbraith. Force

Characteristics of Centrifugal Blood Pump Impellers. in World Congress on Medical

Physics and Biomedical Engineering (WC2003). 2003. Sydney, Australia.

Timms, D. L. and Tan, A. C. C. Performance Characteristics of a Centrifugal

LVAD in Annual Queensland Health and Medical Scientific Meeting, 2003.

Brisbane, Australia.

1.6.4 Patents

Provisional Patent Continuance Application “Fluid Pump” (AU2004/906579), 2004

International Patent Application “Fluid Pump” (PCT/WO2004098677); Inventor –

Daniel Timms, Entering National Phase in International Countries, 2004

Chapter 2 – Literature Review

2-1

Chapter 2

Literature Review

This section provides a focused literature review on many of the important aspects of

implantable ventricular assist device technology relevant to this thesis. Additional

reviewed material relating to the field of mechanical cardiac assistance is included in

Appendix A.

This first section highlights the increasing incidence of heart failure in the

developing world. Current treatment options are discussed, with particular emphasis

placed on the potential benefits of mechanical assistance.

The design requirements of such devices are then outlined, focusing on compatibility

and performance with respect to size, lifetime, blood damage and efficiency. The

notion of device control with respect to patient activity, and the incidence of

ventricular collapse, is then introduced.

A brief review of flow visualisation literature is provided prior to a summary of the

conducted background study in this field. The purpose was to identify regions of

turbulence and stagnation that may generate blood trauma in a centrifugal pump.

Right heart failure can sometimes accompany left heart failure, and the prevalence

and cause of this dysfunction is therefore detailed. This information lends support to

the need for bi-ventricular assistance, with the current device techniques used to

address this situation presented.

Current mechanical assist devices are then briefly reviewed, from the first generation

pulsatile devices, to the third generation rotary devices. Focus is drawn toward the

centrifugal type, with the methods of impeller suspension and rotation presented.

The magnetic bearing technique used for impeller suspension is expanded in the final

section. Current configurations employed by various centres are included, before a

brief summary of the findings from a background study is presented. The purpose of

this study was to identify the various magnetic bearing options available for

complete impeller suspension in each of the VAD designs described in this thesis.


2-2

2.1 Heart Failure and Treatment Options

Incidence of Heart Failure

The incidence of Chronic or Congestive Heart failure (CHF) is increasing each year.

Three million US patients experienced CHF in 1995, with 400 000 new cases

diagnosed each year (McCarthy 1995; Nakazawa, Takami et al. 1997). By 2001,

these figures increased to 4.7 – 5 million and 550 000 respectively, while the

worldwide figure was estimated at 800 000 – 1 million new cases per annum (Data

1996; DeBakey 2000; American_Heart_Association 2001; Anderson 2001; Rose,

Gelijns et al. 2001). Heart Failure also affects infants, with 15,000 babies born each

year with congenital defects (Chen, Smith et al. 1998).

CHF remains a leading cause of morbidity and mortality in the United States,

forming the principle cause of death in approximately 55,000 to 60,000 patients

annually, and a contributing factor in another 200,000-250,000 (Schoeb, Barletta et

al. 2000). Fifty percent of patients diagnosed with CHF will die within one year, and

seventy percent by five years (DeBakey 2000).

The New York Heart Association classifies the Congestive Heart Failure patients

into Class 1 (35%), Class 2 (35%), Class 3 (25%) and Class 4 (5%). Only patients

listed as class 4 are eligible for cardiac transplant. As many as 4000-4500 American

patients from class 4 are listed for transplant each year (Anderson 2001). However,

due to the shortage of donor hearts (less than 3000 worldwide (Rose, Gelijns et al.

2001)), the transplantation rate according to the United Network for Organ Sharing

(UNOS) scientific registry was between 2100-2400 per year in the late 1990’s

(McCarthy 1995; Data 2000; American_Heart_Association 2001), with 770 patients

succumbing to their illness while on the waiting list (Nakazawa, Takami et al. 1997).

In 1995, sixty-six percent of registrants had been waiting over 6 months on the

UNOS list and the average waiting time for a donor organ has now increased to over

300 days for outpatients, contributing to the 15 to 20 percent mortality of waiting

heart transplant candidates (Takami, Otsuka et al. 1998).


2-3

In Australia, cardiovascular disease is the leading cause of death over any other

disease. Currently around 44 per cent (50,797) of all deaths each year are directly

related to heart disease. Of these, 2500 Australians died in 1998 as a direct result of

heart failure ((AIHW) 2001). During this year, 72 heart transplant operations were

performed, with this number ranging from 65-100 from 1994 to 1999. A total 1146

transplants have been performed in Australian and New Zealand from 1984-1999

with the waiting period for transplant between 1-1687 days (Waters 1998; Davies

and Senes 2003). Success rates are high for these patients, with upwards of 80% of

patients still alive after 5 years.

A number of sources have estimated the number of patients that would benefit from

heart transplantation surgery, or ultimately mechanical assistance. The American

Heart Association estimate that annually the lives of between 20,000 and 40,000

Americans would improve with heart transplantation, while Schoeb et al. (2000)

predict between 70 000 – 120 000 and Chen et al (1998) approximate 35,000 to

165,000. McCarthy (1995) and Guy (1998) predict that 35,000 - 70,000 US patients

with Class IV heart failure could benefit from mechanical support, while Anderson

(2001) estimates this number at 100,000 patients worldwide.

Alternative Treatment Options to Heart Transplant

Since a large percent of heart failure is attributed to left ventricle failure, an

implantable LVAD appears as a promising alternative to cardiac transplantation.

Other alternatives include Drugs, Xeno-transplantation (Lewis and Graham 1995),

Cardiomyoplasty (McCarthy 1995) and Stem Cells. However some of these

techniques are currently either ineffective for the entire patient population, or have

ethical issues. For example, although developing pharmacological therapies

improves survival, symptomatology, and quality of life for Class II and III heart

failure, some patients become tolerant or develop side effects that prevent drug use,

and therefore mortality remains high (Anderson 2001).


2-4

Mechanical Assistance

The shortage of donor hearts, and expense of heart transplantation, has accelerated

the development of mechanical cardiac assist devices. Therefore, the demand for

mechanical circulatory assistance is expected to increase significantly as the

incidence of heart failure increases.

Devices can be employed to bridge a patient to heart transplant, to recovery, or as

destination therapy. With the success of mechanical assistance, indications are

widening to include patients not considered transplant candidates.

Such a device will help to alleviate the demand on worldwide health care system by

offering an alternative to expensive drug therapy associated with organ transplant

and other conventional heart failure treatment. Furthermore, the additional capacity

is provided for terminal or waiting list patients to relocate from intensive care units

to cheaper wards, or ultimately leave the hospital and return home.

Bridge to Transplantation (BTT) The survival rate of heart transplantation after the use of cardiac prothesis as bridge

to transplantation is in the range of 90 percent, similar to straight heart

transplantation. However before this stage, 30 – 40 percent of patients die during

prothesis support. Therefore, at least 50 percent of patients on cardiac prothesis

support survive after transplantation. Although this percentage is low, Nose et al

(2000) note that many of these patients would die without mechanical support.

Bridge to Recovery (BTR) It has been reported that diseased hearts would recover if given rest for one year. By

supporting cardio myopathic patients mechanically, it is believed outcomes should

improve substantially, and by allowing the diseased heart to remain, the patient is not

totally device dependent (Nose, Yoshikawa et al. 2000). Some cardio myopathic

patients have been weaned after three weeks of support, while chronic unloading also

improves ventricular function (Takami, Otsuka et al. 1998).


2-5

REMATCH Trial The Randomized Evaluation of Mechanical Assistance for the Treatment of

Congestive Heart Failure (REMATCH) was a multi-centre study supported by the

American National Heart, Lung and Blood Institute. The 3 year (May 1998 – June

2001) randomized trial compared the long term implantation of the TCI Heartmate

VE LVAD (68 patients) with optimal medical management (61 patients) for patients

with end stage heart failure who require, but do not qualify to receive cardiac

transplant (Rose, Gelijns et al. 2001).

The primary objective was to determine the effect of the LVAD on mortality from all

causes, while secondary objectives were to analyse cardiovascular-related mortality

and cases of worsening heart failure, assess functional status and health related

quality of life and finally cost effectiveness.

Due to the risks and lifestyle implications of living with an LVAD, it was concluded

that optimal medical management would provide a more rigorous test for the device.

To qualify for the REMATCH trial, patients must have had class IV heart failure.

They were also ineligible for cardiac transplant for one or more of the following

reasons: Older than 65 years, Diabetes with end organ damage, Chronic renal

failure, Major physical or psychiatric co-morbidity (Cancer, Obesity etc)

25% of the optimal medical therapy group survived at one year versus 50% of the

LVAD group. The survival was 8% and 23% after two years respectively.

The cause of death in the control group was heart failure, whereas sepsis and device

failure accounted for the majority of deaths in the device group. Adverse events, such

as device malfunction, bleeding and infection occurred 2.35 times more often in the

device group. The probability of infection was 28% within 3 months of implantation,

with the device failing 35% of the time after 2 years.

Although the trial indicated an increase in quality of life for patients receiving LVAD

therapy, it is unclear if the magnitude will be sufficient to convince the medical

community that this expensive technology can be reasonably considered an

alternative to the biological replacement of a heart.


2-6

2.2 Ventricular Assist Device Design Considerations

There are many recommendations and requirements to consider when designing an

implantable VAD. For example, the device should be power efficient, avoiding

on/off systems. Atraumatic and anti-thrombogenic characteristics are essential, aided

by the elimination of valves. The device must be durable for long term support and

inexpensive. Additionally, the pump should output pulsatile flow in conjunction

with the failing heart, due to the flow characteristics of a rotary pump (Nose,

Yoshikawa et al. 2000). The service life for a VAD intended for alternative to

transplant use is 10-15 years. Power consumption should be less than 10W with a

maximum temperature of 42OC for device to tissue interface. Ample and

straightforward flow paths should be provided to minimise hæmolysis and

thrombosis formation. Maher et al (2001) concede that the challenge of pumping

blood with undue damage has largely been solved, leaving power system and

management, speed control issues and patient interface issues.

Lifespan / Mechanical Failure

The operational lifetime of a VAD is dependant on the mechanical stability and

reliability of components. Components that move and contact each other are

subjected to stresses and strains that will result in their eventual failure.

Diaphragm type VAD’s are already routinely used as BTT devices, however the

large size and mechanical wear (due to valves) limits their operational life to 2-3

years in patients over 40kg (Schoeb, Barletta et al. 2000; Burke, Burke et al. 2001).

Achievement of a service life of 10-15 years may not be possible with mechanical,

blood immersed bearings, nor rotating mechanical seals. Therefore many groups

have investigated and are currently developing rotary blood pumps with levitated

rotors (Loree, Bourque et al. 2001).

Nose et al (1999) noted that recovery of dilating cardio myopathic hearts can occur

within 6 months – 1 year. Therefore a reliable pump able to operate in excess of 2

years is desired; a characteristic not afforded by axial type pumps supported by

wearable bearing systems due to their relatively high rotational speed.


2-7

Thrombosis

Another concerning factor for implantable blood pump design is thrombosis. Clots

form in areas of stagnation and on foreign bodies or rough surfaces. The potential

for more damage can occur if this clot becomes dislodged (embolism), and carried

away to a smaller vessel, fatally cutting its blood supply (e.g. stroke).

Both Miyazoe et al. (1999) and Nishida et al. (1999) identify the gap beneath the

impeller as the most obvious area for low flow, as blood leaks back from the volute

to this region and has no mechanism to dissipate.

The ability to control the clotting of blood is a feature of human circulation. The

endothelial cells of arteries and veins secrete a variety of chemical substances that

inhibit blood clotting on their walls. Artificial materials used in implants do not have

this ability; therefore the blood may clot when presented with a foreign surface.

This may be overcome with specialized materials or by the use of drugs such as

heparin. Kopp et al. (2002) suggest that bonding heparin to biomaterial surfaces

decreases protein activation.

In the gyro series pump, a pivot bearing supports the impeller. This bearing is the

site for stagnant flow and is the most potential thrombogenic area inside the pump.

In vivo testing performed by Nakata et al. (2000) found firmly attached thrombi at

the top and bottom bearing areas. It is proposed that these thrombi do not dislodge

and form embolic events.

The Heartmate VAD series employ a textured surface to promote the growth of a

protein layer on the blood contacting surfaces (Maher, Butler et al. 2001). This

technique aims to reduce levels of thombosis formation and cell damage by

attempting to recreate the arterial lining.

Hæmolysis

Hæmolysis results from the shearing and rupturing of red blood cells and consequent

release of haemoglobin into the bloodstream. The degree of damage is proportional

to the magnitude of shear force and cell residence or exposure time (Paul, Apel et al.

2003). Many groups have attempted to correlate these factors with actual haemolysis

levels, with most agreeing that a threshold exists sparking significant damage, as

opposed to cumulative damage.


2-8

Threshold

Nevrail (1969) concluded a threshold value of shear stress exists at 1500 dyne/cm2

(150 Pa) for the start of RBC lysis, while the majority of RBC hæmolysis occurs at

shear stress values exceeding 3000 dyne/cm2 (300 Pa). No appreciable damage is

detected when exposure levels were limited to 1000 dyne/cm2 (100 Pa).

Paul et al. (2003) used a coette device to evaluate haemolysis within shear ranges

from 30 – 450 Pas, with residence times of 25 – 1250ms. The results indicate a

threshold exists for elevated haemolysis above 425 Pas and 620ms.

Allaire et al. (1999) indicate shear stress levels in all areas of the pump should be

kept below 500-1800 Pa at all times.

Index of Haemolysis

Hæmolysis is measured with a normalised index (Equation [ 2-1 ]) and standard

procedure outlined in (ASTM_F1841-97 1998).

TQ100

100

Ht-100 V Hb free NIH

××××∆=

[ 2-1 ]

Where NIH ∆freeHb

V Ht Q T

Normalised Index of Haemolysis (g/dL) Increase in Plasma Free Haemoglobin (g/L) Circuit Volume (L) Percent Hematocrit (%) Flow Rate (L/min) Sampling Time (min)

Lewis and Graham (1995) identified that hæmolysis levels are only considered

clinically significant if plasma free haemoglobin values average more than 0.01

g/100L. Values of NIH for different centrifugal type blood pumps are given in Table

2-1.

Table 2-1 - The index of hæmolysis of blood pumps

Continuous Flow Pumps Index of Hæmolysis (g/100L) BP-80 Nikisso

Kyoto MSCP CFVAD#2 CFVAD#3

Baylor C1E3 Baylor PI601 Baylor PI701 Baylor KP101

Ventrassist

0.0017 0.00066 0.015 0.0086 0.0124 0.0007 0.0028 0.004 0.01 0.005


2-9

Clearance Gaps

Nishida et al. (1999) propose the design factors greatly affecting the level of

hæmolysis are the radial and axial clearance gaps between the rotating impeller and

stationary pump casing. These gaps typically range from 0.2mm - 3mm, and should

be optimised to produce the lowest levels of hæmolysis whilst maintaining the

highest possible hydraulic efficiency. The rotational speed of the impeller directly

affects the level of shear stress encountered in the clearance gaps. A larger peripheral

velocity results in larger shear stresses, therefore the rotational speed of the impeller

should be minimised.

In a study on the Ventrassist device conducted by James et al. (2003), axial clearance

was found not to influence the level of haemolysis. The clearance gaps were small

(0.07 – 0.3mm), and although CFD analysis predicted an increase in regional shear

stress, blood cell exclusion and a smaller residence time in these gaps accounted for

the maintenance of low haemolysis levels. Despite previous claims that impeller

surface roughness has no effect on haemolysis, the group produced results indicating

a reduction in NIH for hand polished (0.1-0.2 um) surfaces compared to machine

finished (0.2-0.7um).

Many authors present results of increased haemolysis with decreasing radial

clearance below 0.5mm. However, it is suggested that further reductions in clearance

gaps below 0.1mm may actually improve haemolysis results. The majority of red

blood cells are thought to avoid entering the gap, and when combined with a smaller

fluidic volume and shorter residence time, fewer cells are damaged (James,

Wilkinson et al. 2003).

Turbulence

Another aim for hæmolysis reduction is smooth flow. Turbulent flow increases

stresses on the blood and although they may not exceed the rupture stress for RBC’s,

undetectable trauma can occur leading to a decrease in the haemolytic threshold of

the whole blood (Allaire, Wood et al. 1999). This is likely to occur at the outlet

region, by blood colliding with casing wall (Nishida, Asztalos et al. 1999).


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Heat Generation

Another form of blood damage involves exposure to excessive heat levels, which

cause proteins in the blood to denature. Proteins are sensitive to their environmental

conditions, and therefore irreversible damage occurs if blood is exposed to heat

levels higher than 44oC (Tortora 1987). Heat may be generated at points of contact

between moving components as well as electrical power sources.

Yamazaki et al. (1998) used a purge fluid system to keep the shaft seal below 40

degrees in a device operating at 2500rpm and producing 7 L/min at 100mmHg

consuming 9 Watts of power.

Tasai et al. (1994) conducted heat generation tests on a polycarbonate motor housing,

which demonstrated excessive temperature rise. The material was replaced with

anodised aluminium, which exhibits excellent heat conductivity. An effective

thermal path was established to distribute the motor heat flux to the circulating

blood. At worst case, the heat flux from the device was calculated at an acceptable

level to surrounding tissue, i.e. less than 0.062 W/cm2

Song et al. (2004) included a thermal analysis in a CFD model of the CFVAD4. The

study concluded that the temperature rise due to magnetoelectric heat generation and

thermal dissipation in surrounding tissue and blood was insignificant.

Takatami et al. (1998) evaluated heat generation in the miniaturised Gyro PI pump.

The study revealed the temperature inside the actuator rose to 43.5 deg when

operating at 6.16 Watts.

Wakisaka et al. (1998) presumed that heat absorption of the surrounding tissue was

about 0.01-0.02 W/cm2. They observed thermal adaptive responses of the lung and

muscle tissue under a heat load of 0.02-0.06 W/cm2. In these cases, the capillaries in

the surrounding smooth fibrous tissue might play an important role in cooling down

the pump. Particularly in smaller devices, since the heat absorption capacity of the

surrounding tissue is decreased by reducing the pump surface area.

Size / Anatomical Compatibility

Current commercially available VADs are large and heavy, limiting their application

to large adults. The desirable outcome from a new VAD design would be the ability

to implant in patients of all size. Rotary type VAD’s have the potential to achieve


2-11

this goal, since the geometry is intrinsically smaller than their pulsatile counterparts.

Device size is a concern for successful implantation. It must not only conform to the

available space in the patients’ abdomen, but a reduced size is essential to

accommodate small children and women.

Physiological Control

Physiological control of ventricular assist devices is an important feature to provide

sufficient perfusion to patients with changing levels of physical activity.

Wu et al. (2004) proposed a control algorithm that uses the change in pressure across

the device and motor signals as feedback signals, obtained from long-term reliable

sensors. A safe range is set for the end diastolic ventricular pressure (-3 to 15mmHg)

to prevent ventricular suction as well as pulmonary congestion. The ultimate goal of

the controller was to return perfusion, mean arterial pressure and end diastolic

pressure to within the normal range. Maintenance of these values while transitioning

from conditions of rest, sleep and exercise for variable levels of ventricular function

was demonstrated within a few heartbeats.

Ventricular Collapse

Olegario et al. (2003) performed a study aimed at controlling the outflow of two

centrifugal pumps acting as a total artificial heart. Sudden and large changes in

outflow from rotary pumps produce negative inflow pressures which can increase the

incidence of atrial or ventricular wall suction, particularly in the left heart. This

predicament can be alleviated by reducing the pump speed which reduces outflow.

This technique is effective for LVAD situations, as the failing natural ventricles

remain and can assist limited flow. However, rotary pump TAH devices rely solely

on the device to produce flow, and thus any reduction is pump speed is met with

critically low perfusion. Since the left pump operation is based on the total peripheral

resistance (1/R control), the proposed solution for the latter case involves the control

of right pump outflow to maintain sufficient left atrial blood volume. That is, left

atrial pressure is monitored, and any reduction is compensated by an increase in right

pump outflow. However, the delay between the right pump and left atrium may result

in an excessive increase in right outflow causing right atrial suction. Therefore the

pulmonary circuit dynamics must be integrated into the control algorithm.


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2.3 Flow Visualisation

Flow visualisation is an important technique used to highlight areas where

modifications may improve the fluid flow pattern. A specialised application of flow

visualisation is in the area of blood pump design (Ikeda, Yamane et al. 1996; Allaire,

Wood et al. 1999).

Consideration must be given to the traumatic effect of the pump on blood. The

incidence of haemolysis and thrombosis are most damaging to red blood cells. The

former results from shearing and rupturing of red blood cells (RBC) and release of

haemoglobin into the bloodstream (Yamane, Clarke et al. 1999), while the latter

results from the formation of blood clots in regions of stagnant or low flow (Allaire,

Wood et al. 1999). By employing flow visualisation techniques, it is possible to

identify areas of high shear, turbulence and low flow or stagnation. These studies are

often conducted in conjunction with computational fluid dynamics (CFD) to refine

the model. Results can be categorised as either qualitative or quantitative.

Qualitative

A qualitative study can instantly identify regions of low flow or stagnation as well as

turbulent regions by analysing captured images using high speed videography

(Nishida, Asztalos et al. 1999). Another technique uses a film of high viscosity oil

painted on the surfaces of interest (Ichikawa, Nonaka et al. 2002). This provides

information on flow patterns at the fluid contacting surfaces only. Qualitative

information is useful for identifying the regions of the pump susceptible to

thrombosis formation and blood cell destruction respectively. Investigators may then

target these regions for further quantitative analysis.

Quantitative

Quantitative analyses commonly use particle image velocimetry (PIV) techniques to

identify particle velocities (Tsukiya, Taenaka et al. 2002). This data can be used to

reproduce the boundary layers within the pump, providing an opportunity to

calculate shear stress levels. The oil film technique can translate quantitative values

of shear stress by incorporating the oil viscosity; however the margin of error is

relatively large. Quantitative information can be related directly to thresholds of

shear stress estimated to cause haemolysis (Nevrail 1969; Allaire, Wood et al. 1999).


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Flow Visualisation Study

A flow visualisation investigation (Timms, Tan et al. 2003; Tan, Timms et al. 2004)

was undertaken as background research on a scaled up centrifugal pump. The pump

geometry did not reflect the final VAD designs; however important qualitative

observations relevant to all centrifugal blood pump designs were made, which

influenced these final VAD designs.

The study was conducted in three stages using the particle seeding and high speed

videography method;

1. Preparation of a rig for testing to be conducted,

2. Utilisation of adequate equipment and implementation of an experimental

technique to produce results, and

3. Qualitative analysis of captured data to identify target regions for further

detailed quantitative analysis.

Images were captured from identified regions of interest at design and below design

flow rates. It is envisaged that the final implanted VAD will not always operate at

design conditions as the level of assistance required is dependant on the individual

patient’s level of heart failure. Therefore, turbulence and stagnation at off design

conditions should not be overlooked.

The flow visualization technique described provided valuable insight into the internal

flow patterns within the centrifugal blood pump design, identifying the region

beneath the impeller as a stagnation zone together with the outflow and clearance

regions as potential sites for haemolysis.

Stagnation beneath the impeller prompted investigations into the elimination or

improvement of washout flow in this region. Techniques such as washout holes

would improve flow, while implementing a double sided impeller would eliminate

the problem completely.

Recirculation was found at the cutwater section. This was further emphasized when

the pump was operated at capacities below the design value, as encountered in the

practical application of the implanted VAD.


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2.4 Right Heart Failure and Bi-Ventricular Support

The main goal of mechanical support is to increase the blood flow to end organs.

Since the left and right hearts are in series, an increase in flow to the systematic

circulation, as observed in LVAD support, will result in an increase in venous return

to the right heart. Therefore the cardiac output of the right heart must increase.

Clinically, if the left heart output is doubled with the institution of a LVAD, the

output of the right ventricle must match (Goldstein and Oz 2000).

Causes of Right Heart Failure

Goldstein and Oz (2000) explain that the introduction of LVAD support may unmask

pre-existing right ventricular dysfunction, by attempting to increase the venous return

beyond the capacity of the right ventricle. This right ventricular dysfunction may

only become apparent after the left ventricle is assisted. To obviate the need for Bi-

or R-VAD assistance, patients receive treatments that attempt to offset existing or

developing pathological conditions.

Goldstein and Oz (2000) continue that with LVAD support, unloading of left

ventricle reduces its volume and pressure. This may result in the interventricular

septum bulging to the left during right ventricle systole, thus reducing the efficiency

of the right ventricle contraction. However, many studies of normal hearts conceded

a change in right ventricle geometry did not affect overall performance. Therefore, it

is apparent that anatomic ventricular interactions are not responsible for the profound

right ventricle failure encountered in patients receiving LVAD support, compared

with the effects of pre-existing pathological conditions. For example, the detrimental

effects of ventricle interaction were negligible compared to the effects of ischemia.

An LVAD was unable to improve the impaired right ventricle cardiac output, due to

the inability for the right ventricle to pump sufficiently through the pulmonary

vasculature. The cardiac output was found to improve with the introduction of right

ventricle assistance or Bi-VAD.

Lewis and Graham (1995) supported these factors, citing altered ventricular

interdependence, increased pulmonary vascular resistance and changes in right

ventricle loading as predominant causes for the onset of right heart failure.


2-15

Need for Bi-Ventricular Support

Nose et al. (1999), Yoshikawa et al. (2000) and Lewis and Graham (1995) each note

that approximately 25-40% of patients receiving LVAD support develop right heart

failure syndrome, which can lead to sepsis and/or multi-organ failure. They explain

that multi-organ failure is induced by poor liver portal circulation as a result of high

venous pressures above 20mmHg. This leads to 30 percent of VAD patients

succumbing to this cause of death (Nose, Yoshikawa et al. 2000).

Kavarana et al. (2002) concede that the introduction of inhaled nitric oxide and

phosphodiesterase inhibitors has reduced the use of right ventricular assist devices,

however right ventricular dysfunction is still a serious problem in patients receiving

left ventricular assist devices.

Yoshikawa et al. (2000) suggest that 50 - 60 percent of the LVAD population require

Bi-ventricular assistance. Goldstein and Oz (2000) identified that this support is

dependant on the patient population, with usage varying from 10% to 85% of support

cases. Suffice to say, a large number of the patient population can be salvaged from

multi-organ failure by introducing biventricular support as this reduces the higher

venous pressure (Nose, Yoshikawa et al. 2000).

Farrar et al. (1997) also recognised the need for the introduction of an RVAD, since

in this study, 48% of the patients received Bi-VAD support as a bridge to cardiac

transplantation.

Hence, there is an increasing need for the development of a Bi-Ventricular assist

device and thus current research is focussed on this goal.


2-16

Current Bi-VAD’s and Techniques

Nose et al. (1999) and Nose et al. (2000) outline the current bi-ventricular assistance

technique of implanting two devices (LVAD and RVAD). However the large size of

these devices, particularly the pulsatile types, poses a problem for anatomical

compatibility. Thus they suggest the non-pulsatile rotary blood pump is the only

option.

Current development of a pulsatile type combined Bi-VAD for a total artificial heart

(TAH) centre around extracorporeal and paracorporeal systems, such as the moving

actuator mechanism proposed by Park et al. (2003). However, this device is

relatively large and has a moving actuator which is prone to wear.

To produce the function of a total biventricular assistance with continuous flow

pumps, Nose et al. (1999) implanted two identical VAD’s in the abdominal wall for

left and a right ventricular bypass. Each pump had its own independent control and

actuation system. The study lasted for 72 days and was terminated due to functional

inflow obstruction.

Yoshikawa et al. (2000) then developed a small yet efficient RVAD that could be

used in conjunction with a conventional LVAD to provide Bi-VAD support. The

LVAD operated under performance characteristics demanded by the systemic

circulatory system, whilst the RVAD met the pulmonary circulatory system

requirements.

To this date, Bi-VAD technology is either too large or necessitates having to implant

two independently working pumps. The development of a compact single continuous

flow Bi-VAD would increase the application of this form of assistance to encompass

the smaller body surface area patient population.


2-17

2.5 Current Mechanical Assist Devices

History has witnessed the development of many types of artificial cardiac device.

Guy (1998) reported on the development of pneumatically driven blood sacs, roller

pumps, pusher plate and diaphragm pumps. Maslen et al. (1998) identified these

types as inherently flawed with reliability problems, which led to the development of

devices based on a rotating impeller.

The most common use of continuous flow centrifugal blood pumps is in short term

surgical procedures involving cardio-pulmonary bypass (CPB). Despite this success,

Miller et al. (1990) and Allaire, et al. (1996) suggest that a variety of complications

served to delay their application as viable risk free long-term total cardiac

replacement devices, and as such were viewed only as bridge to transplant devices.

However, Maslen et al. (1998) predicted that with improved design and testing, these

devices could function as a permanent replacement for heart failure.

Types of Mechanical Assist Devices

The classification of mechanical assist devices relates to their characteristic outflow,

followed by their generation or bearing configuration).

Pulsatile

Franzier et al. (1992) reported that although pulsatile type pumps are similar in

operation to the human heart, they are relatively large and require valves which can

cause blood cell trauma.

First Generation The first generation of pumps incorporate sacs, diaphragms or pusher plates to

produce pulsatile outflow from the device. They are relatively large and are powered

either pneumatically or electrically; with a maximum expected lifetime of two or

three years.

Non-Pulsatile / Continuous Flow

The continuous flow devices generally incorporate an axial or centrifugal type rotary

impeller. Kawahito et al. (1996) outlined that although axial flow pumps are

relatively simple in their design and construction, they require high rotating speeds

and are usually supported by contact pivot bearings. Sasaki et al. (1992) and Ohara et

al. (1994) describe that centrifugal flow pumps have high efficiency, but initial


2-18

designs required frequent replacement due to moving component wear and heat

generation around the impeller shaft, seals and bearings. This imposed a significant

limitation if they were to be used for long-term assistance. Several techniques have

been proposed to solve this problem, ranging from complete magnetic-suspension

(Akamatsu, Nakazeki et al. 1992), two-pivot bearing support (Ohara, Makinouchi et

al. 1994), free impeller suspension (Ohara and Nose 1994), and finally completely

passive hydrodynamically suspension (Watterson, Woodard et al. 2000).

Second Generation Second generation pumps are classified as continuous flow axial or centrifugal

pumps with impellers supported by blood immersed or pivot bearings, and are rated

to five year durability. The majority of axial flow blood pumps are classified in this

category.

Third Generation Magnetically levitated and hydrodynamic non-contact impeller suspension

characterise the third and latest generation of blood pumps. A ten year lifetime is

expected due to the elimination of component wear from mechanical contact

(Takatani 2001).

Table 2-2 categorises some of the current mechanical assist designs.

Table 2-2 – Current Mechanical Assist Devices

TAH Abiocor (Stevenson and Kormos 2000) Cardiowest - Jarvik 7 (Arabia, Copeland et al. 1999; Copeland, Smith et al. 2001) Penn State / 3M (Stevenson and Kormos 2000)

Pulsatile LVAD

Novacor (Stevenson and Kormos 2000; Copeland, Smith et al. 2001)TCI Heartmate I (McCarthy 1995; McCarthy, Schmitt et al. 1996)Thoratec IVAD Arrow Lionheart

Axial Flow LVAD

Heartmate II (Burke, Burke et al. 2001; Maher, Butler et al. 2001) Jarvik 2000 (Jarvik 1995)Debakey/Micromed (Tayama, Olsen et al. 1999; Stevenson and Kormos 2000)

Centrifugal LVAD

Heartmate III (Loree, Bourque et al. 2001; Maher, Butler et al. 2001)Heartquest(CFVAD) (Khanwilkar, Olsen et al. 1996; Bearnson, Olsen et al. 1998)Ventrassist (Watterson, Woodard et al. 2000)Ibaraki (Masuzawa, Kita et al. 2000; Masuzawa, Kita et al. 2001)Baylor GYRO series (Yoshikawa, Nakata et al. 1999; Yoshikawa, Nonaka et al. 2000)


2-19

Axial VAD Song et al. (2003) provided a comprehensive review of all axial flow devices. Most

devices, such as the HEARTMATE II, JAVIK 2000 and Micromed DeBakey, are

generally limited to second generation pumps; however, the INCOR device

incorporates third generation magnetic technology. Axial type devices are

characteristically small, and consequently incorporate rotary blades that spin at

10,000 to 20,000 rpm, which may lead to excessive shear stress. Additionally, these

pumps require stationary diffuser guide vanes, providing a site prone to thrombus

formation. Apart form the INCOR device, the service life of axial flow pumps is

predicted at below five years, due to the contact pivot bearing support mechanism.

Centrifugal VAD Continuous centrifugal flow devices address many of the inherent problems

associated with first generation positive displacement mechanical assist devices.

Advantages

The characteristic advantages over first generation devices include a lack of valves,

no compliance volume, anti-thrombogenecity due to a smaller blood contact area,

mechanical simplicity, smaller size, higher efficiency, lower costs, fewer moving

elements, and lower hæmolysis (Park, Nishimura et al. 1996; Nishida, Uesugi et al.

1997; Tsukiya, Akamatsu et al. 1997; Takami, Otsuka et al. 1998; Takatani 2001).

Miller et al. (1990) declare the superiority of centrifugal pumps due to the ability to

generate large flow rates and pressure heads, and the ability to easily power and

control such systems electrically. In addition, large ventricular volumes, valves or

diaphragms are non-existent, ensuing increased durability with less moving parts.

Yamazaki et al. (1997) support the use on the basis that rotational speeds are lower

than axial pumps, which would result in longer bearing life. The elimination of

stationary blades also reduces the risk of thrombosis formation.

Akamatsu et al. (1992) and Curtis et al. (1999) indicate the advantages of operational

simplicity, compactness, and low cost compared with other mechanical assist

devices.

Finally, Smith et al. (1999) point out the ability to change flow and head output by

simply changing the RPM of the impeller is beneficial in accommodating body

physiological changes.


2-20

Disadvantages

Akamatsu et al. (1992) were aware of centrifugal pump limitations, stating that high

friction in the bearing and seal generates a large amount of heat. Yamazaki et al.

(1997) warned that if this heat becomes too high, the proteins within the blood will

denature resulting in irreversible blood damage.

Wampler et al. (1999) described how the process of sealing the blood at the

driveshaft severely limited the durability of initial rotary blood pumps while Aztalos

et al. (1996) expressed the concern that excessive rotational speeds lead to large

shear stresses in the radial clearance.

Finally, Goldstein and Oz (2000) outlined the possibility for regurgitant flows should

the pump stop. Since no valves are employed, blood flow back into the heart could

possibly worsen the patient’s heart failure. Vandenberghe et al. (2002) assessed the

effect of rotary blood pump failure on the hemodynamic condition using a

mathematical cardiovascular model. The simulated results indicated an increase in

aortic flow to compensate for the regurgitant flow through the device. This was

achieved by an increase in stroke volume and contractility to maintain perfusion and

arterial pressure. The study concluded that the acute hemodynamic state was not

lethal, and indeed reversible should support resume within a 15 minute timeframe.

Methods of Impeller Support and Rotation Allaire et al. (1996) acknowledged that the simplest way to suspend a rotating pump

impeller is via a shaft, bearing and seal system. Unfortunately, they point out that this

arrangement introduces unreliability, as seals are subject to abrasive failure and

provide a site for heat generation, blood stagnation and clotting. Yamane et al.

(1995) thus predicted the operational lifetime of this pump is limited to a few days.

Wampler et al. (1999) proposed that until a viable alternative to the shaft seal is

implemented, the evaluation of this type of pump for use as a LVAD is restricted.

A number of authors have proposed various methods of suspending the centrifugal

blood pump impeller. Sipin et al. (1997) used a purged ball bearing system, Maslen

et al. (1996) developed a bearing/seal system, Watterson et al. (2000) utilized

hydrodynamic bearing forces and Yamane et al. (1997) implemented a monopivot

bearing. However, applying a magnetic motor-bearing system, potentially

minimizes all problems associated with the aforementioned forms of suspension

(Masuzawa, Ezoe et al. 2003).


2-21

Pivot / Blood Immersed Bearings

Blood immersed bearings form another means of impeller support. This type of

bearing may be surface lubricated with blood, and allow rubbing contact of ball

bearings. Thompson et al. (2003) proposed a system that incorporated wear proof

UHMWPE roller bearings with a service life in excess of ten years. Thrombosis

complications are avoided by purging the bearing system with saline at a rate of

30mL/hour.

Hydrodynamic Bearings

Alternatively, a thin fluid film between rotor and housing can be employed to

provide true hydrodynamic support with no rubbing contact (Wampler, Lancisi et al.

1999; Tansley, Cook et al. 2000; Tansley, Vidakovic et al. 2000; Watterson,

Woodard et al. 2000). However, Allaire et al. (1996) and Maslen et al. (1998) warn

that the bearing clearances may cause shear damage to red blood cells and lead to

hæmolysis.

Magnetic Bearings

Maslen et al. (1998) detail a promising method of impeller suspension employing

magnetic fields to levitate the impeller. Magnetic bearings boast non-contact

suspension with large clearance gaps, thus eliminating any point-to-point contact to

reduce shear stress. These systems are described in further detail in the next section.


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2.6 Design, Performance and Force Characteristics

Introduction

A review of literature relating to centrifugal blood pump design, performance and

hydraulic force characteristics is provided to outline current research and identify

areas for further investigation.

The literature review targeting the design of centrifugal blood pump impellers

provides a summary of the design parameters suggested and selected by various

research groups. Little information is provided on the design of volute casings,

diffusers and inlet configurations with respect to performance, although one recent

study used CFD techniques for assessment of these configurations (Song, Wood et

al. 2004). Few studies have focussed on the effect of volute clearance gaps on blood

damage.

The difficulty in predicting the performance characteristics of a centrifugal blood

pump due to their inherent dissimilarities to industrial applications was outlined in

the next section of this review. A set of dimensionless pump curves was sourced as a

means to compare new pump prototypes. Finally, conflicting reports were identified

regarding the effect of axial clearance gap above open impeller blades on pump

performance.

Relatively few studies feature in the final review of this section on centrifugal blood

pump hydraulic forces, revealing an important area for further investigation.

Knowledge of these forces is particularly important when designing a magnetic

bearing to completely suspend the impeller. To add to the review, a number of

hydraulic force studies conducted on industrial pumps were included. Both

experimental techniques and results are provided to help select the most appropriate

method of force evaluation, as well as provide a comparison to the data presented

later.


2-23

Centrifugal Blood Pump Impeller Design

Roto dynamic Impeller Design

Qian et al. (2002) identified four important factors for impeller design in centrifugal

LVAD’s; Dimension (Pump Type and Size), Vane Form (Profile) including exit

angle and vane number. Curtas et al. (2002) also identify vane number and profile as

critical, but also suggest that the shroud configuration and vane exit width are

important parameters for performance.

Dimension / Pump Type The pump’s dimension dictates the choice of pump type, and is selected according to

pump speed, pressure and flow rate (specific speed) at the design point. Pump type

selection based on specific speed calculations reveals the notion that centrifugal

pumps are more suitable than axial flow pumps for the rotary blood pump

application. Comparing the normalised head and flow curves for both centrifugal and

axial flow pumps of similar size shows the necessity for the axial flow pump

impeller to rotate at 140% greater tip speed than that of a centrifugal design. For

example, the practical range for impeller tip speed of an axial and centrifugal pump

for the LVAD application is between 10-15m/s and 4-7m/s respectively (Yamane,

Ikeda et al. 1996). Haemolysis levels are mostly affected by the creation of shear

stresses at the impeller tip, which are affected by the impeller rotational speed, but

more closely related to the impeller tip speed. This implies that a centrifugal pump

would produce a lower level of haemolysis.

Tansley et al. (2000) used the pump design techniques of Stepanoff (Stepanoff 1957)

to dimension/size their centrifugal pump for the LVAD application. This analysis

concluded the pump would operate at a specific speed (Ns) of 1000, leading to an

impeller diameter of 40mm rotated at 2500rpm. An impeller tip speed of 4.9 m/s was

calculated after a speed constant of 0.95 was assumed.

Vane Profile, Exit Angle

Curtas et al. (2002) suggest that an optimum vane discharge angle exists for each

centrifugal pump application. Additionally, designs should consider the effects of the

inlet blade angle. Occasionally, inlet blade angle equals outlet blade angle, for which

case the profile should follow that of a logarithmic spiral. However, it may be

desirable to vary the blade angle as a function of radius. CFD focussing on

variations on blade profile between common entrance and discharge angles revealed


2-24

little effect of alternate profiles (logarithmic or otherwise) on performance.

Qian et al. (2002) make stronger claims that vane profile should take the form of a

logarithmic spiral, with a discharge angle (β2) of 30º. Their conclusion is founded on

the prediction that the fluid will follow streamlines based on mass, motion and

energy conservation. With an impeller following these principles, no turbulence

would be apparent in the vane channels and blood damage would be minimal. These

assertions were confirmed with favourable performance and haemolysis test results.

A study performed by Takano et al. (2000) analysed the effect of vane angle on the

performance of a centrifugal blood pump. The experimental results of a 35mm

diameter impeller showed that the radial vanes (β2 = 83º) produced the desired

pumping characteristics at a lower rotational speed (2900 rpm) to that required by an

impeller incorporating slightly backward facing (β2 = 73º) vanes (3280 rpm).

However, hydraulic efficiency was not compared in this study.

These assertions are confirmed by CFD results presented by Curtas et al. (2002), who

concluded that the most head was produced from straight radial blades as opposed to

backwardly curved vanes. Efficiency was considered in this study and found that

although a greater head was produced, it was at the expense of efficiency.

Yamane et al. (1996) supported these claims with an indication that the slope of the

head-flow curve is determined by the discharge angle (β2).

However, Smith et al. (1999) indicate that caution must be given to designs

incorporating high impeller inherent pressure/flow curves (produced from high vane

discharge angles) as they are characterised by low efficiencies which may impact on

battery size/life and the degree of heat generation. They therefore recommend that

the design of an impeller to achieve a desirable pressure/flow relationship an

efficient manner is a valuable exercise.

Finally, Masuzawa et al. (1999) conducted a flow visualisation study on the effect of

backward, radial and forward facing blades for the blood pump application.

Exit Width

Curtas et al. (2002) employed CFD techniques to determine the effect of exit blade

height (with a common axial clearance) on performance. The explanation centred on

the fact that a larger exit width (3mm) led to a reduced meridional velocity (cm2) and


2-25

thus an increase in head for a given rotational speed when compared to a smaller exit

width (1mm). The implementation of a larger blade width departed from

conventional design guidelines which called for a smaller width to maintain an exit

to inlet area ratio (a2/a1) closer to unity. They suggest the deviation in results from

conventional guidelines is acceptable since the effect of this area ratio on small

pumps is not well documented, and many other design parameters fall outside the

normal regions.

As mentioned, the exit width affects the magnitude of meridional velocity. This

directly impacts on the ratio (λ) of absolute tangential velocity (cu2) to this

meridional velocity (cm2), which is maintained between 2 and 7 in industrial pumps.

Therefore, a disproportionately large exit width may lead to an excessively small

meridional velocity, which in turn increases the value of λ above 7, and as high as

28 (Tansley, Vidakovic et al. 2000).

Vane Number The number of vanes contained on an impeller affects the level of slip, or blade

recirculation. For example, a large blade number reduces slip as there is less space in

the vane channel to allow recirculation. However, other fluid effective losses are

dominant in small pumps with low flow and speed. Therefore viscous effects

dominate the core flow within impeller passages, while velocities are generally not

sufficient to produce separation from the blades (Curtas, Wood et al. 2002).

Li (2000) conducted Laser Doppler Velocimeter experiments that confirmed the

effect of fluid viscosity on pump performance. As expected, higher viscosity fluids

result in a rapid increase in disc friction losses over the impeller shroud and hub, as

well as hydraulic losses in the flow channels. This suggests that a larger blade

number might result in a drop in efficiency, due to viscous drag on the larger surface

area. Furthermore, if slip is less of a concern, larger blade discharge angles may be

permissible to produce more head without the anticipated losses.

In the study performed by Qian et al. (2002), vane number and angle was selected as

a result of haemolysis testing performed on impeller configurations incorporating 5-7

vanes at discharge angles between 20-40º. The results revealed six vanes at a

discharge angle of 30º are kindest with regards to haemolysis.


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Shroud Configuration Many of the current rotary blood pump impellers incorporate an impeller hub and

front shroud to improve hydraulic efficiency (Anderson, Wood et al. 2000; Chua,

Ong et al. 2003; Masuzawa, Ezoe et al. 2003). However, concerns are raised with

regard to the level of haemolysis or potential for thrombus formation within the

clearance gaps, particularly beneath the hub.

Chua et al. (2003) investigated velocity profiles within the axial clearance gap of the

shrouded type impeller of the magnetically suspended centrifugal blood pump

(MSCBP), while Anderson et al. (2000) used CFD techniques to investigate the

recirculation from impeller outlet to inlet within the axial clearance gaps. This

information is useful in predicting pump efficiency and effect on haemolysis or

thrombosis.

The effect of semi-open and open type impellers on hydraulic performance was

investigated by Tsukiya et al. (2002), revealing a slight drop in efficiency with the

former type. This result was due to an increase in recirculation flow beneath the

shroud to the impeller inlet via the included washout holes, caused by a larger

pressure gradient along the flow path.

Roto dynamic Inlet and Outlet Design

Few published studies have targeted the impact of inlet and outlet configurations on

centrifugal type rotary blood pump designs. However, Song et al. (2004) conducted

a CFD study to evaluate these inlet and outlet devices.

Inlet devices (Straight inlet, Inlet Volute, Inlet/Flat Elbow) shown in Figure 2-1

were compared with regard to anatomical compatibility, performance and predicted

impact on haemolysis. The straight area-converging inlet, although best from a

turbo-machinery standpoint, would require an effective pipe length of 250 mm to

provide straight, uniform and unsteady flow into the impeller eye. Obviously, this is

not practical in terms of implantable rotary blood pumps. The inlet volute is adopted

in most industrial single stage pumps. Although boasting improved anatomical

compatibility for implantable blood pumps, the main drawback (inlet swirling)

creates a challenge in impeller design to prevent this affecting overall pump

performance. Additionally, high shear stresses were found to occur in this

configuration, indicating the potential for elevated levels of haemolysis. Results from


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the CFD analysis revealed the pump should incorporate a conventional inlet elbow,

which is hydraulically equivalent to a straight inlet for the low specific speeds

associated with rotary blood pumps. Further investigations revealed advantages of a

flat type inlet elbow incorporating a streamlined baffle on the further side, since

circumferential velocity is minimised and swirls are almost completely removed.

Figure 2-1 – Inlet (Left) and Outlet (Right) Devices

(Song, Wood et al. 2004)

The relevant outlet devices discussed by Song et al. (2004) and displayed in Figure

2-1, centre on the use of a single type volute and outlet diffuser. The cross sectional

area of this volute increases from cutwater to outlet at an angle corresponding to the

direction of the absolute velocity vector at impeller discharge. The exit volute design

is suggested to initiate with the calculation of a throat area by assuming a value of

throat velocity obtained from theoretical references. Reference was made to the

importance of cutwater diameter selection to improve performance and reduce

recirculation between the cutwater and impeller. The final step involved drawing the

expanding base circle from cutwater to throat before implementing a 6-10 degree

divergence of the outlet diffuser. Figure 2-2 describes the double volute

configuration implemented by Chua et al. (2003) in the magnetically suspended

centrifugal blood pump (MSCBP).

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2-28

Figure 2-2 –Double Volute Configuration

(Chua, Ong et al. 2003)

Rotary Blood Pump Performance

Smith et al. (2004) noted the difficulty in predicting the performance of rotary blood

pumps from previous industrial pump data. Conventional pump designer’s benefit

from the collection of correlated empirical data combined with extensive practical

experience. Unfortunately this practice is not apparent in rotary blood pump design

due to the infantile stage of the field coupled with inherent dissimilarities from

industrial applications.

Smith investigated and compared numerous published pump performance results,

producing a family of rotary pumps developed for cardiac assistance (Figure 2-3).

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Figure 2-3 - Non-Dimensional Performance of Rotary Blood Pumps

This graph shows the dimensionless performance curves for a number of axial, mixed flow and centrifugal blood pump designs, which can be used to compare new prototype designs

(Smith, Allaire et al. 2004)

Three non-dimensional grouping ratios (flow, head and power coefficient) were used

to compare different designs (Table 2-3). This data provides the ability to compare

new designs to a family of rotary blood pumps. For example, a shut-off head

coefficient of 0.5 - 0.6 was suggested acceptable for the blood pump application.

Table 2-3 – Dimensionless Coefficients

(Smith, Allaire et al. 2004)

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2-30

The transition of Reynolds number from laminar to turbulent regime was noted at the

conventional 5.0 E5 instead of 2.3 E2 as suggested in cases of tube flow. Blood

pumps in general were found to exhibit a low Reynolds number (approx 7.6 E4),

which negatively influences pump scaling. As such, difficulties are presented when

attempting to use old pump design data to predict the performance of a new blood

pump. Furthermore, high fluid viscosity and the need for small size, high efficiency

and good flow patterns, necessitate a pump that falls outside the conventional design.

Pump performance is directly affected by hydraulic losses. These were suggested to

be a combination between viscous (windage, friction) and non-viscous (blade

thickness, mixing). The hydraulic efficiency of the axial flow blood pumps

investigated did not reach the heights of industrial pumps of the same specific speed,

whereas the centrifugal types were more comparable. This further emphasises the

mismatch of the axial type pump to the cardiac assist application.

Centrifugal pumps incorporating an open or semi-open impeller are also subjected to

losses associated with leakage over the vanes. This was emphasised by Stepanoff

(1957), whereby the magnitude of leakage was dependant on the ratio of impeller

vane height to axial clearance. However, Takano et al. (2000) obtained results

suggesting hydraulic performance and thus leakage loss above the 4mm high

impeller vanes did not increase with a variable axial clearance gap of 1.5 to 5.5 mm

in the Gyro series VAD. This is contrary to results from industrial pumps; however

they attributed the discrepancy to the difference in specific speed and flow

magnitude. Further investigations revealed little difference in haemolysis for the

axial clearance variations, and suggested that the radial clearance gap is much more

predominant in effecting haemolysis levels.


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Centrifugal Impeller Forces

Centrifugal Blood Pump

The force characteristics of a centrifugal blood pump impeller are important in

determining the stiffness capacity of the bearing system responsible for impeller

suspension. Resultant thrust is of particular importance when employing magnetic

bearings to support the rotor, as precise loads must be specified in order to complete

the bearing design (Japikse, Marscher et al. 1997).

The impeller should expect to encounter loads resulting from static and dynamic

hydraulic forces within the pump housing, impact forces from everyday activities

such as walking, and impeller gravitational force.

Impact forces do not usually not exceed 2-3g in everyday life (Tansley, Cook et al.

2000)). Accelerations of 4 - 6g for more than a few seconds may cause blackouts

and death, while decelerations of between 50-75 g will cause fatalities in humans.

Gerhart et al. (2002) performed drop tests of the CorAide VAD to simulate

survivable automotive crash data. Results indicated no change in device performance

following simulated falls from 1.2m and 4m which produced decelerations of 12 -

15g and 35 - 44g in all axes respectively.

The Ventrassist device can withstand dynamic accelerations up to 100 m/s2 (10g)

before impeller touchdown (Tansley, Cook et al. 2000). Impact loads are further

catered for by an exotic coating on impeller surfaces which allow minimal

touchdown without pump cessation. A CFD study of the hydrodynamic force on the

open type Ventrassist impeller, with varying axial rotor position, was performed by

Qian and Bertram (2000). Predicted radial forces of 0.15-0.3N acted at an angle

between 78 and 90 degrees (toward/parallel cutwater). Axial lift forces were

computed between 0.2 and 1.4N.

A study on radial forces encountered by a centrifugal blood pump impeller Allaire et

al. (1996) determined a maximum resultant radial force of 1.5 N when the axis of

impeller rotation was parallel with the horizontal. Since the rotor static weight was

3.5 N, the force reduction was attributed to the centring effect of water in the

clearance gaps which offset the impeller weight. This phenomenon is experienced

by industrial pumps employing axial flow seals (labyrinth) which produce centring

forces, positive stiffness and damping effects (Allaire, Kim et al. 1996).


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Axial force was the topic for investigation by Takami et al. (1997) on the GYRO

centrifugal blood pump. This pump employs a permanent magnetic coupling

beneath a semi-open vane type impeller and thus a magnetic attraction force exists in

the negative axial direction. This force is offset by the floating hydraulic force in the

positive axial direction, created as a result of asymmetric pressure distribution within

the pump cavity. This force was estimated to range from 10 – 40 N when operating

the 65mm diameter impeller at 1500 – 3000 rpm respectively. The actual force was

30N for the cardiopulmonary bypass application (300mmHg) and 10N for ventricular

assistance (100mmHg). These results were used to determine the magnetic coupling

distance, which alters the magnetic attraction force (Takami, Makinouchi et al. 1997)

Axial force was also investigated in a CFD study performed by Curtas et al. (2002).

Centrifugal Industrial Pump

Stepanoff (1957) performed a set of simple radial and axial force measurements in a

centrifugal pump based on impeller geometry, operating head and normalised

capacity. Although comprehensive, this study concentrated on the forces experienced

by the impeller of a centrifugal pump designed for one specific speed. Realising this

shortcoming, Agostinelli et al. (1960) modified the experiments to account for the

effect on radial force of pumps designed for various specific speeds.

Flack et al. (1984) provided a comprehensive literature review on the mechanisms,

measurement and prediction of both static and dynamic radial thrust, while Lorett et

al. (1986) targeted the interaction of impeller and volute, particularly operating at off

design capacity. Guelich (1987) then provided an overview of the source of static

and dynamic forces and techniques to measure them.

Both Adkins et al. (1988) and Jery (1985) examined the effect of a predetermined

amount of static impeller whirl on static and dynamic force production.

Recently Baun et al. (2003) conducted a series of experiments aimed at determining

the effect of various impeller and volute combinations on radial thrust force, with

parallel examination of hydraulic performance. They also analysed the effect of a

statically eccentric impeller located within a circular volute, revealing an optimum

offset position may be obtained to improve hydraulic efficiency and reduce hydraulic

force to levels comparable to conventional spiral volute pumps (Baun, Kostner et al.

2000).


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The following sections describe the experimental techniques and results obtained

from a selection of these studies.

Experimental Techniques

Pressure Tappings

Radial thrust can be measured by recording values of pressure about the impeller

circumference by a number of tappings, and integrating the pressure profile (Japikse,

Marscher et al. 1997).

This method requires simple equipment, however the measurement and integration is

cumbersome. Additionally, the accuracy of this technique is low, since the pressure

measured is at the volute wall only, and is integrated only over the impeller outlet

width. Therefore, the contribution of pressure acting on the shroud and hub faces is

neglected and the resultant forces is generally concluded lower than the actual value

(Guelich, Jud et al. 1987).

Non-uniformities in fluid momentum around the impeller are neglected, since simple

pressure tappings do not account for changes in fluid velocity within the volute.

Resultant radial force calculated using this technique is generally lower than the

actual force encountered by the impeller. To address this issue and improve

accuracy, Pitot-static probes may be inserted into the volute. Static pressure and

velocity profile within the volute can be measured, which when integrated, reveals

static pressure and momentum force (Flack and Allaire 1984).

Bearing Stresses

Forces transmitted from the impeller to the casing are measured by strain gauges or

load cells equipped to brackets supporting the bearings. The setup is calibrated by a

known mechanical imbalance and/or statically loading in the impeller plane. This

technique is the most common method used for force evaluation (Stepanoff 1957;

Uchida, Imaichi et al. 1971; Flack and Allaire 1984; Guelich, Jud et al. 1987; Adkins

and Brennen 1988; Japikse, Marscher et al. 1997).

This method does not reveal static and dynamic contributions to overall force,

instead the resultant of ALL forces acting on the rotor is measured, including the

labyrinth forces (Guelich, Jud et al. 1987).


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Shaft Stresses

Strain gauges may be placed directly on the rotating shaft to measure the resulting

radial forces. Results are similar to those provided by the bearing technique; however

the initial setup is more complex (Flack and Allaire 1984; Guelich, Jud et al. 1987).

Shaft Deflection

Proximity probes may be used to detect the deflection of the shaft in response to

increased impeller force. This method is simple, quick and not much modification is

needed to an existing pump to install the probes. However, accuracy is moderate due

to the variations in bearing clearances (Guelich, Jud et al. 1987).

Magnetic Bearing Load Cells

Supporting the shaft of the centrifugal pump with magnetic bearing creates the

ability to measure impeller forces. By recording the power requirements of the radial

and axial magnetic bearing, the force characteristics may be inferred (Japikse,

Marscher et al. 1997; Baun and Flack 1999). This technique requires a high

magnetic bearing stiffness to reduce any eccentricity caused by changing forces.

Literature Results The theoretical force characteristic trends described previously are confirmed by

experimentation (Stepanoff 1957; Lazarkiewicz and Troskolanski 1965; Lorett and

Gopalakrishnan 1986; Baun, Kostner et al. 2000; Karassik, Messina et al. 2000;

Baun and Flack 2003).

For example, Baun et al. (2003) produced radial thrust results for four and five vane

impellers in single, double and concentric volutes operating over various flow

capacities. Figure 2-4 presents a normalised capacity versus non-dimensional

measured force. The latter corresponds to the value of thrust factor “K” in the

empirical equations used to predict radial force.


2-35

Figure 2-4 – Experimental Measurement of Radial Thrust Magnitude and Direction

A slight print error was observed in the original diagram. The correct symbolisation requires the exchange of filled and unfilled symbols representing force characteristic and direction.

(Baun and Flack 2003)

The same group also discovered that thrust characteristics achieved in a spiral volute

casing were replicated in a concentric casing while operating an impeller at a certain

offset distance (Baun, Kostner et al. 2000). Figure 2-5 describes the results obtained

from that study. An impeller with no offset created force characteristics typical of a

concentric casing, while an offset in the general direction of the cutwater (35o)

increasingly reproduced spiral volute characteristics. This revealed that radial thrust

is affected by the relative static eccentricity of the impeller within a circular casing.

Figure 2-5 – Values of Radial Force Factor (K) for Eccentrically Located Impellers in a

Concentric Casing

(Baun, Kostner et al. 2000)

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2.7 Magnetic Bearings for Rotary Blood Pumps

Maslen et al. (1998) identified a promising method of impeller suspension that

utilises magnetic fields to levitate the impeller in a high clearance envelope.

Wampler et al.(1999) recognise that if the magnetic field is active, complete

suspension of the rotating impeller is possible. Maslen et al. (1998) further declared

that a range of diagnostic information on impeller position and active suspension

force may be extracted.

Magnetic bearing technology boasts non-contact suspension, eliminating any point-

to-point contact. Allaire et al. (1996) and Maslen et al. (1998) realised this technique

provided a mechanism for reduced wear and friction, thus simultaneously reducing

the potential for cell mechanical trauma, degree of generated heat, and sites for blood

stagnation.

Karassik and McGuire (1998) state that bearing power is only needed for magnetic

excitation, and is generally lower than lubrication system associated losses.

Incorporating permanent magnets into the magnetic bearing system (PM biasing),

while using active electromagnets for fine control can further reduce power

consumption.

Chen et al. (1998) also support that bearing configurations consisting of two passive

radial bearings and one active magnetic thrust bearing are most attractive in terms of

reduced power consumption, simplicity and reliability. However, they concede the

physical limits and corresponding performance consequences must be addressed

when integrating the bearing system into the pump. Such limits are apparent when

dealing with disturbance forces which require the bearing to have sufficient strength

and stiffness to keep the rotor levitated. These forces may arise from shock transients

(cough), rotor weight, and hydraulic forces. Some systems attempt to use DC

brushless motor designs to provide rotation. Conventionally, these motors produce

large permanent magnet/stator forces in the axial or radial directions, which are

handled by conventional bearings. However these forces are reduced by the

requirement for larger clearance gaps to allow for blood flow (Xu, Wang et al. 1997).

Ultimately, an increase in the bearing capacity leads to increased magnetic bearing

size. Wampler et al. (1999) suggest that the electromagnets in active magnetic

bearings are generally large and heavy, and require proximity sensors fed back to


2-37

servo circuits. An increased complexity results, and therefore may compromise

reliability. Therefore Schweitzer (1994) insist a comprehensive knowledge of

mechanics, electronics and control is required to fully understand their operation.

Bartha et al. (1999) realised that to provide stability in an active bearing control

system, the precise feedback of rotor positional information along each axis is

essential. One proposed method is self-sensing (sensorless) magnetic bearings,

which need no external position sensors (Maslen and Noh 1996). The position

information is deduced from the air gap dependent properties of the electromagnets.

The magnetic motor-bearing system is required to keep the impeller centred within

the pump clearances whilst providing rotational torque / actuation along all six co-

ordinates (3 translational x,y,z and 3 moments θx,θy,θz). Although all such actuation

is provided magnetically, Maslen et al. (1998) indicate that one co-ordinate is not

fixed and requires a motoring mechanism. This is substantially different from the

bearing system used in the remaining 5 co-ordinates. Developments in magnetic

bearing technology applied to centrifugal VAD’s have concentrated on implementing

a combination of passive and active magnets to completely suspend the impeller.

Passive Magnetic Bearing

Passive magnetic bearings incorporate an arrangement of permanent magnets for

suspension. Depending on their configuration, passive permanent magnets can be

used as either attracting or repelling radial or axial magnetic bearings. However

Schweitzer et al. (1994) reveal that it is not possible to obtain a stable equilibrium,

and at least one unstable co-ordinate remains. This is the basis of Earnshaw’s

theorem; “It is impossible to maintain a body at a given position in free space with

only magnetostatic forces”.

Yamane et al. (1995) suggested that one of the impeller’s six degrees of freedom can

be supported mechanically; the other five are supported with permanent magnets. In

device shown in Figure 2-6, a “top” shaped pivot mechanically supports the axial (z)

degree of freedom. The pivot is located at the bottom centre of the impeller,

surrounded by four permanent magnets coupled through the housing to the DC

motor. The radial degrees of freedom are supported by repelling permanent magnets.

The system can work stably without special sensors or controllers due to the

simplicity of permanent magnets.


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Figure 2-6 – Monopivot Magnetic Suspension Pump

(Yamane, Nishida et al. 1997)

Chen et al. (1998) orientate two axially polarized permanent magnet rings so the

stationary ring attracts the rotating ring of the rotor (Figure 2-7). Therefore, if the

rings are radially misaligned, a restoring force (reluctance centring effect)

proportional to the misalignment is generated.

Figure 2-7 – Radial Restoration Using Permanent Magnets

(Chen, Smith et al. 1998)

Khanwilkar et al. (1996) and Chen et al. (1998) believe the incorporation of a passive

permanent magnet bearing into the system may alleviate the power requirements of

the active magnetic bearing to successfully suspend the pump impeller.

Arrangement of the permanent magnets to counter the generated forces in the

directions of highest magnitude is essential to optimise the bearing efficiency.

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Active Magnetic Bearings

Bearings involving active reluctance or lorenz magnetic forces take the form of

radial (x,y), axial (z) or total (x,y,z,θx,θy) suspension, relating to the number of

degrees of freedom (DOF) actively supported. The radial type bearing is thinner, as

the magnetic mechanism is located in the circumferential space. An axial bearing

improves performance due to the ability to incorporate more magnetic material in the

unused axial surface area of the impeller (Masuzawa, Ezoe et al. 2003).

Radial Type Radial type magnetic motor bearings employ active magnetic forces to suspend the

rotor in the radial (x,y) direction and electromagnetic coupling to provide rotational

torque to the impeller.

Masuzawa et al. (2000) proposed a radial motor-bearing to the reduction in the

controlled number of degrees of freedom (Figure 2-8).

Figure 2-8 – Radial Magnetic Motor-Bearing

(Masuzawa, Onuma et al. 2002)

The radial motor-bearing requires placement of the electromagnets in the same plane

as the rotor magnets, allowing a thinner overall shape. These electromagnets are

responsible for controlling two radial directions (x,y), and the desired impeller

rotation about the third axis (θz). To simplify control of the remaining degrees of

freedom, the axial movement (z) and tilting moments (θx,θy) are restricted by passive

stability (Figure 2-9). This design was improved by Masuzawa et al. (2002) using

FEA to optimise the stator. The study also assessed the effect of double and single

volute configurations, concluding that a double volute reduced the radial

displacement and thus radial bearing power from 1.3 to 0.7 Watts.

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Figure 2-9 – Principle of Passive Restoring Force

Okada et al. (1997) also created a magnetic/motor bearing that levitates and controls

the impeller in the two radial degrees of freedom (x,y). Rotation about the z-axis (θz)

is achieved by electromagnets surrounding the circumference of the impeller. A

digital signal processor provides the control of levitation and rotation in this system.

Gap sensors provide feedback on the rotor’s position, which is used to calculate coil

current from the summation of motor and levitation control currents.

Axial Type Axial type magnetic bearings employ active magnetic forces to suspend the impeller

in the axial direction (z). These forces can also help to counteract the gyroscopic

coupling moments (θx,θy). The remaining DOF are stabilized via other techniques.

Okada et al. (1997) produced a magnetic bearing that controls one degree of freedom

(z). Rotation is controlled by the phase of the magnetic flux between the rotor and

stator, while changing the amplitude of the magnetic flux controls levitation. The

remaining four DOF are passively restricted. Rotation and levitation are controlled

using a digital signal processor (DSP). A gap sensor produces a displacement signal,

which is used to determine the current required for levitation, and added to the

rotational current. The levitation control algorithm is based on a simple proportional

derivative (PD) controller. Only one degree of active control is achieved in this

configuration, and the total stability of levitation was poor. The efficiency of this

arrangement is low, as the motion control requires passive stabilisation of the

remaining four DOF. The main advantage is the relative ease of control

implementation.

Restoring Force

Restoring Force Rotor

Rotor

Stator

Stator

Axial Displacement

Tilt Displacement


2-41

Masuzawa et al. (2003) are developing a magnetically suspended centrifugal blood

pump with an axially levitated motor (Figure 2-10).

Figure 2-10 – Axial Magnetic Motor-Bearing

(Masuzawa, Ezoe et al. 2003)

The impeller is actively controlled in the axial and tilting degrees of freedom, while

relying on passive stability for restriction of radial displacement. Design conditions

demand 6 Watts of power for the motor bearing, assuming hydraulic efficiency of 30

percent while operating at a rotational speed of 2000 rpm and 0.03 Nm torque. To

reduce the magnitude of negative stiffness, the total air gap between top and bottom

stators and the rotor is 3mm. To improve pump hydraulic efficiency, the impeller

and casing axial clearance is reduced to 0.4mm, while the radial volute gap is

1.5mm. Displacement while operating the pump in a mock circuit indicated

sufficient bearing capacity; however bottom impeller touchdown was evident at

increased rotational speeds. The bottom stator was responsible for rotation, while the

top stator controlled tilt only. Therefore, the impeller migrated toward the bottom

stator as torque demand increased, until touchdown resulted. Although radial

movement was contained within the given clearance, a significant movement of 0.6

mm was observed. Further studies of volute pressure distribution were recommended

to assess radial passive stability.

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Total Total active magnetic bearings utilize electromagnetic forces to support the impeller

in all five stabilized DOF and rotate the impeller about the sixth. The University of

Utah and Virginia have developed a completely levitated and rotated centrifugal type

pump (Allaire, Kim et al. 1996; Allaire, Kim et al. 1996). However, Masuzawa et al.

(2000) warn that this requires a sophisticated control mechanism to control all DOF.

Allaire et al. (1996) successfully designed and constructed a prototype incorporating

a five axis active magnetic bearing system. An active control system is used with an

analogue PID controller to prevent the rotor from contacting the pump housing in the

axial direction. An eddy current sensor provides the rotor position feedback to the

controller. Magnetic bearings have the capability to provide large impeller/housing

clearances in the pump (0.3 – 1.0 mm), which help to increase flow rates and

minimize induced shear stresses. Furthermore, active control is capable of limiting

vibrations within the range of 0.03 - 0.05 mm. The radial bearing at outlet has a 16

pole, planar, radial bearing stator.

In the model constructed by Maslen et al. (1998), a circular array of horseshoe

magnets is directed at the inlet face of the impeller. They account for three axial co-

ordinates, one translational (+vez) and two moments (θx,θy). An 8-pole radial stator

with tapered faces directed at the discharge edge of the impeller account for two

translational co-ordinates (+ve/-ve x,y) as well as –ve z. This array of magnets includes

9 electromagnets pre-biased to simplify control.

Motor When properly selecting a driving motor and its control, the torque characteristics of

the machine to be driven must be considered.

Wampler et al. (1999) transmitted torque via a brushless axial flux gap motor

coupled to the impeller. Permanent magnets embedded in the impeller are rotated by

an electromagnetic field produced by a flat motor stator, commutated using back-

EMF sensing. The flat motor helps to reduce overall size, not to mention minimizing

the air gap. Maslen et al. (1998) used an array of neodymium-iron-boron permanent

magnets embedded into the underside body of the impeller. An associated slotless

10-pole, 3-phase stator embedded in the pump casing provides the magnetic coupling

and driving mechanism at a relatively high efficiency of 75 percent.


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Magnetic Bearing Investigation

As indicated in the review, numerous magnetic bearing configurations are under

development for use in the implantable VAD application. These generally feature

unique characteristics suitable to the geometry of each device.

A number of configurations were therefore considered for potential implementation

in the VAD designs proposed in this thesis. These ranged from radial type magnetic

motor-bearings with outer and inner rotors, to axial type self bearing motors.

The potential success of each VAD design is dependant on the understanding and

correct selection of magnetic motor-bearing types.

Preliminary investigations were conducted on an 8-pole slotless and 4-pole slotted

outer rotor (Kim, Abe et al. 2003) , as well as an 8 pole slotted and 8 pole slotless

inner rotor (Tokumoto, Timms et al. 2002). All configurations were of the Lorenz

force type. Reluctance type bearings were not studied in detail.

A brief summary of this study is outlined below, with further details provided in

Appendix F. A comparison of the operating characteristics of each self bearing

motor investigated is presented in Table 2-4.

Table 2-4 – Comparison of Self Bearing Motor Operations Characteristics

Inner Rotor Outer Rotor

SLOTLESS SLOT SLOTLESS SLOT

MAX RPM 5500 2100 - 10,000 MAX EFFICIENCY 89.3% 58% - -

CURRENT-FORCE STIFFNESS 3.5 N/A 11.5 N/A 1.2 N/A 10 N/A MAX TORQUE 0.17 (Nm/A) 0.49(Nm/A) - -

The investigation revealed that slotless inner rotor type self bearing motors achieve

stable levitation at higher rotational speeds than slotted types. However, slotted

types are characterised by extensively higher current-force stiffness.

The maximum speed of the slotted inner rotor type (2100rpm) is insufficient for

VAD operation. Depending on the impeller geometry, the self bearing motor should

achieve stable levitation at speeds up to 3000rpm. This may be overcome by

improving the magnetic circuit and permanent magnet design to reduce the

destabilising cogging torque.


2-44

Electrical current stiffness in the all slotless types is considered insufficient to

prevent impeller touchdown in the implantable VAD application, especially with the

existing rotor weight. Incorporating a slotted type self bearing motor would allow for

greater bearing capacity (approx 10N/A) to counter hydraulic pressure loads and

external shocks encountered during expected operation. For example, if the

maximum device supply voltage is 5V and the coil resistance is 1 Ohm, the bearing

could generate forces of 50N. This would accommodate a shock of 50G for a 100g

rotor. Torque generation was more than sufficient for both slotted and slot type inner

rotor motor-bearings.

Concerns were raised during the investigation over the passive axial stability of these

self bearing motors, especially in the slotter outer rotor configuration. The lack of

passive axial stiffness generated from the radial active control would not appear

sufficient to accommodate expected hydraulic axial thrust forces generated during

conventional single sided impeller pump operation (Timms, Tan et al. 2004).

Conclusion

A number of magnetic motor-bearing configurations were researched and/or

experimentally investigated to indicate their potential for use in the proposed Bi-

LVAD and Bi-VAD designs.

The slotted type bearings produced a larger suspension capacity, while the slotless

inner rotor bearing could achieve a higher rotational speed than its slotted

counterpart, due to the elimination of destabilising cogging torque. All designs

demonstrated weak passive axial stability.


2-45

2.8 Review Summary

The ageing population of the developed world is accelerating the incidence of heart

failure. The most successful treatment of end stage heart failure, the heart transplant,

is limited by the availability of donor organs. Alternative treatments are under

development; however a successful mechanical assist device would provide a

virtually unlimited supply of universally compatible replacement organs. Previous

trials have demonstrated the benefit of VAD treatment over optimal drug therapy,

however device failure and large physical size reduces their availability to the

general patient population.

Many of the essential design issues for the development of ventricular assist devices

were outlined in the review. The potential for blood trauma must be considered

carefully when attempting to decide on pump features. Regions of stagnation, high

shear and high temperature must be eliminated or confined to acceptable levels.

Mechanical durability is of equal importance if the device implantation period is to

approach the expected lifetime of a transplanted heart. Furthermore, this goal must

be achieved with an anatomically compatible design that is suitable for implantation

into the majority of the general CHF population.

The background flow visualisation study was able to identify regions of recirculation

and stagnation within a scaled up centrifugal blood pump. These results, coupled

with an identified cause for thrombosis found in the literature, prompted the

investigation of a double suction impeller for each VAD application to reduce

stagnation beneath the impeller, as well as highlight the need for careful volute

design for smooth outflow transition.

The current bi-ventricular assistance technique using two devices (LVAD and

RVAD) is limited by the large size of these (usually) pulsatile extracorporeal

devices. The development of an implantable compact single continuous flow Bi-

VAD would increase the application of this form of assistance to encompass the

smaller body surface area patient population. The review supports the aim of Bi-

VAD development, as many studies concluded between 17-60% of patients receiving

LVAD support develop right heart failure. This occurrence can result in up to 30% of

patients succumbing to sepsis and/or multi-organ failure due to high venous

pressures. This onset of RHF after LVAD support was commonly attributed to pre-


2-46

existing pathological conditions. Many authors suggest that this therapy may unmask

right ventricular dysfunction, alter ventricular interdependence, increase pulmonary

vascular resistance and change right ventricle loading. Anatomic ventricular

interactions due to septal shift were not considered responsible for the profound right

ventricle failure encountered in these patients. The introduction of right ventricle

assistance and ultimately bi-ventricular assistance successfully improved cardiac

output is such cases.

Mechanical cardiac devices are classified according to their characteristic outflow

(pulsatile or non-pulsatile), followed by the type of bearings employed (first, second,

third generation). First generation pulsatile devices were identified as relatively large

with a mechanical lifetime below three years. The non-pulsatile devices were

questioned in regard to their continuos flow output, as well as potential for

regurgitation upon pump failure. However their smaller size and longer lifetime has

accelerated these devices to the forefront of current VAD research. These devices

are generally axial or centrifugal in configuration. The smaller size of the axial type

was recognised, however the requirement for higher rotational speeds and stationary

guide vanes were suggested to promote excessive blood trauma. Furthermore, the

typical second generation pivot suspension limits the service life to below five years,

which is not favourable for the application of destination therapy. The centrifugal

pump was commonly recommended as the most suitable type for long term cardiac

assistance, due to its lower rotational speed, higher efficiency and reduction of blood

trauma creating features. Furthermore, these pumps could take advantage of the latest

third generation bearing technology to completely suspend the impeller using

hydrodynamic or magnetic bearing forces. This reduces the number of moving

components, potentially increasing device lifetime to beyond ten years.

The critical components for centrifugal blood pump design were identified in this

review. The importance of pump type and size revealed the suitability of the lower

specific speed centrifugal pump over the axial flow counterpart. The performance of

the centrifugal impeller was found to produce a larger pressure head at lower

rotational speeds for higher vane discharge angles, albeit at the expense of efficiency.

Forward facing vanes were investigated for flow patterns, but not performance.

Larger exit vane widths were also found to increase values of pressure head for a

given rotational speed, due to a reduction of meridional velocity. This


2-47

recommendation departs from conventional industrial pump design, however is

common in the investigated blood pump impeller designs. The effect of slip and

viscous drag should be considered when determining the number of impeller vanes to

be included between the commonly employed impeller shroud and hub. A higher

vane number reduces slip, while a lower number reduces viscous drag, both working

with the shroud and hub to improve efficiency. Centrifugal blood pump inlet and

outlet devices were reviewed. An inlet elbow was described as hydraulically

equivalent to a straight inlet for the low specific speeds; however a flat type inlet

elbow may provide improved anatomical compatibility. Finally, the calculation of

throat area by assuming a value of throat velocity obtained from theoretical

references was identified as the critical parameter for exit volute design.

Since the performance characteristics required of an implantable VAD fall outside

conventional pump design, it is difficult to use old industrial pump data to predict the

performance of a new VAD prototype. Furthermore, there is no large collection of

empirical VAD data, so all new prototypes must be tested. However, from the

collection amassed by one author, a shut off head of 0.5-0.6 is expected for the

LVAD application. Mismatch of the axial type VAD to the application of ventricular

support is emphasised by their inability to match the efficiency of the centrifugal

type. Finally, a discrepancy is apparent between two authors regarding the effects of

varying the axial clearance above the semi-open type impeller vanes, with and

without a change in vane height, on centrifugal pump performance. One conforms to

theory suggesting that efficiency and thus performance would drop in all centrifugal

pumps with an increasing clearance. However the other author discovered no

alteration in performance. This result is not consistent with theory, and therefore

formed the basis for further investigation.

Resultant hydraulic thrust is an important consideration for the design of third

generation VAD’s that relies on the generation of magnetic or hydrodynamic forces

for complete impeller suspension. Values of hydraulic thrust may be calculated or

obtained experimentally. The prediction of radial thrust relies on experimentally

determined thrust factors for use in calculations. Values of these specific speed

dependant factors are provided by a number of authors. However, they are generally

based on much larger industrial pumps. Since there is limited information available

regarding the practical measurement of hydraulic thrust on centrifugal blood pump


2-48

impellers, the requirement for impeller force measurement in a centrifugal blood

pump application was identified. Such an investigation will provide experimental

results that can be compared to predictions made using each authors’ thrust factors.

This would identify the most suitable set of thrust factors, and enable the accurate

prediction of hydraulic thrust in future centrifugal VAD prototypes. The bearing

stress technique was subsequently pursued due to the relative high accuracy and ease

of set-up compared to the alternative methods. The inability to easily separate the

static and dynamic axial force components was not considered a problem, since the

relative magnitudes are not expected similar in an implantable VAD. The force

results from the literature confirm the predicted theoretical trends of force and

capacity. This information will be useful for comparison and verification of the

results presented in the later investigation.

Magnetic bearing configurations can incorporate active (electromagnetic) or passive

(PM) impeller support. Active support is generally confined to the radial or axial

DOF for simplicity. This technique allows for larger clearances to improve washout,

and active control to increase suspension forces to counter external shocks.

The ability to control the VAD using feedback stimuli is an important feature for

inclusion in the device, and a suitable evaluation environment should be provided for

validation. The VAD should have the ability to respond to changes in the patient

activity level, as well as incidences of over-pumping and ventricular collapse.


2-49

2.9 Conclusion

This review supports the development of new VAD technology as an alternative to

heart transplantation. The Bi-LVAD and Bi-VAD presented in this thesis can

accommodate the cardiac patient population requiring either left ventricular or bi-

ventricular assistance. The Bi-LVAD double impeller would improve magnetic

suspension capacity, and reduce complications of thrombosis. The Bi-VAD device

would provide bi-ventricular assistance to smaller patients, as the single device is

smaller than the present technique of implanting separate LVAD and RVAD pumps.

To achieve the goal of VAD development, a suitable prototype testing facility is

required. A complete mock circulation loop of the systemic and pulmonary

circulation systems is required to test both VAD prototypes. This facility should have

the ability to replicate varying degrees of left and right heart failure, to assess the

pumps ability to return pathological hemodynamics to normal.

Each pump must be designed with careful consideration of the design issues

presented in this review. In order to meet the hemodynamic requirements of the

vascular system and improve pump flow patterns, a procedure must be adapted to

design the impeller and volute components.

The selection of a third generation magnetic bearing technology for impeller

suspension may alleviate many of the problems associated with earlier generation

devices. To assist in the bearing design, precise loads required for impeller

suspension need to be determined. Many of these loads in the clinical setting are

transient (coughing, walking), however fluid pressures created during pump

operation are permanent, and may result in hydraulic thrust. Measurement of this

thrust will help quantify the exact force requirements of the bearing system to

minimise impeller touchdown.

3-1

Chapter 3

Mock Circulation Loop

The requirement for in-vitro evaluation of left- right- and bi- ventricular assist device

performance prompted the design and construction of a new systemic and pulmonary

mock circulation loop.

This chapter presents aspects for the development of the mock circulation loop,

which includes pulsatile left and right ventricles coupled with vascular compliances

and resistances.

A description of the parameters within the natural circulatory system required for

replication is initially provided. An overview of current basic and advanced mock

loops is then included, before a simplified simulation model is presented. This model

assisted in the development of the physical loop, which is detailed in the following

section. Finally, results are presented that verify the loop’s ability to replicate the

hemodynamic conditions of the entire cardiovascular system in response to normal

heart function and varying levels of left and/or right heart failure.

This newly developed circulation loop has the potential to provide valuable

cardiovascular device performance feedback prior to expensive and time intensive

in-vivo animal trials.

Chapter 3 - Mock Circulation Loop

3-2

3.1 Introduction

By reproducing the human circulatory system in-vitro, it is possible to refine VAD

designs by ascertaining their effect on circulation. This accelerates the design

process, and is necessary to comply with FDA regulations before clinical animal and

human invivo trials (Patel, Allaire et al. 2003).

Various mock loops have been constructed by a number of research centres for

testing the viability of their VAD design. These loops range from a basic circulation

loop to advanced loops incorporating variable parameters imitating the physiological

parameters of the human circulatory system.

It is an important criterion for the more advanced mock circulation loops to

reproduce human physiological conditions, both healthy and otherwise, without an

assist device. Inserting the VAD will then evaluate the desired improvement of flow

and pressure, analogous to increased perfusion (Patel, Allaire et al. 2003).

3.2 Background - Description of Physiological Parameters

To develop a mock circulation system, all parameters in the human circulatory

system to be reproduced must be identified and understood.

3.2.1 Circulatory System

The circulatory system encompasses the heart and vascular network. Arteries and

veins form this network, which distributes blood to organ and tissue capillaries

before providing a return to the heart (Figure 3-1). Additionally, a small amount of

blood flow leaving the left heart supplies the bronchial tissue and returns directly to

the left heart, bypassing the pulmonary circulatory system. This results in slightly

variable balance of left/right heart output flow. Highest resistance to flow is

experienced in the arterioles, while most blood volume is stored in veins. The

vessels of the vascular network help adjust the blood flow and pressure via a number

of autoregulatory mechanisms. Particularly, the arteries diminish the pressure pulse

wave created by the intermittent heart ejection, while the veins are capacitance

vessels responsible for regulating the rate of blood returning to the heart (Klabunde

2004).


3-3

Kidneys

Legs

Pelvic Organs

HepaticSplenic

Trunk

Coronary

Arms

Headand Neck

HEART

Lungs Bronchial

ArterialVenous

Figure 3-1 – Circulatory System Figure 3-2 – Circulatory Pressure Distribution

3.2.2 Blood Volume and Pressure Distribution

Blood pressure is highest in the aorta, and slightly reduces in the distributing arteries.

However, most pressure drop (approx 70%) occurs in the arteriole-capillary section

(Klabunde 2004), as displayed in Figure 3-2.

Total blood volume of an average adult human is estimated at 5000-5600 ml

(80mL/Kg). This volume is tightly regulated by complex mechanisms, one of which

is water and sodium excretion by the kidneys. Blood volume significantly effects

circulatory pressure distribution and cardiac output, and therefore plays an important

role in hemodynamic regulation.

Approximately seventy percent of total blood volume is contained in the veins with

the remaining thirty percent in the arteries (Tyberg 2002). Table 3-1 further details

the relative blood volumes (William, Stevens et al. 2003).

Table 3-1 – Blood Volume Distribution

Circulatory Section % of Total Blood Volume Blood Volume Systemic Artery Volume 16 800 ml Systemic Vein Volume 64 3200 ml

Systemic Capillary Volume 4 200 ml Pulmonary System Volume 9 450 ml

Heart Blood Volume 7 350 ml


3-4

3.2.3 Heart Functionality

Cardiac Cycle

The cardiac cycle consists of five phases (Klabunde 2004). Figure 3-3 provides a

view of these phases in relation to pressure distribution and electrical activity.

Figure 3-3 – Phases and Pressure for one Cardiac Cycle

(Klabunde 2004)

PHASE 1 – Ventricular Filling

a. Rapid Ventricular Filling – As the ventricle pressure drops below the atrial

pressure, the A-V valves open and blood re-fills the ventricle.

b. Reduced ventricular filling – As the ventricles fill with blood and expand, the

compliance reduces and pressure increases, thus reducing the flow.

halla



3-5

PHASE 2. - Atrial systole

The atrial muscles contract during this phase to propel blood from the atrium to the

ventricle. This accounts for only 10% of ventricular filling during rest and 40%

during exercise, as most filling occurs passively before the atria contract. Back flow

from the atria to supplying veins is prevented by the blood’s inertia, as well as a

“milking” atrial contraction.

PHASE 3. - Isovolumetric Contraction

The valves of the heart are closed and no flow occurs as the pressure in the ventricle

increases. The volume is the end diastolic volume.

PHASE 4 – Ventricular Ejection

a. Rapid Ejection – When the pressure in the ventricles exceeds the aortic and

pulmonary artery pressures, the aortic and pulmonic valves open and blood is

ejected.

b. Reduced Ejection – Ventricular active tension decreases and the rate of ejection is

reduced. Atrial pressures rise gradually due to filling from venous return.

PHASE 5. - Isovolumetric relaxation

The valves close when the pressure in the ventricles reduces below that in the

arteries, preventing backflow. The pressure of the remaining blood in the ventricle

continues to drop, and the volume is the end systolic volume.


3-6

Heart Rate

Heart rate is the number of cardiac cycles per minute (bpm). Rates of 60, 60 and 120

bpm were chosen for rest, treated heart failure and exercise conditions respectively.

Systolic and Diastolic Period

The Systolic period refers to the ventricular contraction time within each cardiac

cycle. This parameter is also referred to as the percentage of the cardiac cycle in

systole, and can range from 30-50% depending on the condition. Diastolic period

refers to the time within each cardiac cycle that the atrium and ventricle relax. The

ventricular diastolic parameter is defined by the remainder of the cardiac cycle

period after ventricular systole. The percentage time in ventricular systole for use in

calculations within this chapter is assumed at 40% for each condition of rest, exercise

and heart failure.

Ventricular Contractility and Relaxation

Contractility refers to the rate at which the ventricle can develop pressure, while

relaxation describes the rate at which pressure falls within the ventricle after systole.

Contractility varies with heart function, increasing for exercise while decreasing in

cases of heart failure. Relaxation is a function of the passive relaxing of the

ventricular wall, and is assumed to be the same for each heart condition presented.

However, some forms of heart failure can influence this parameter.

3.2.4 Vascular Hemodynamics

Hemodynamics refers to the factors influencing perfusion through the circulatory

system. Organ perfusion is driven by an arterial to venous pressure gradient, which

is dependent on vascular resistance. A relationship between Cardiac Output (CO),

blood pressure (P) and vascular resistance (R) is shown in Equation [ 3-1 ].

RPP

Q veinartery −=

[ 3-1 ]

Since the left and right hearts operate in series, their long term average flow must not

be different. However, due to the varying properties of the circulatory systems, the

short term average may be slightly different allowing sections of the system to hold

different amounts of blood.


3-7

3.2.5 Physiological Pressures

Arterial Blood Pressure

The ejection of blood from the ventricle into the arterial network results in a

characteristic pressure wave. Pulse pressure (PP) and mean arterial pressure (MAP)

are calculated from the values of systolic and diastolic pressure. Typical systolic

(Psys) / diastolic (Pdias) pressures expected in the left and right circulatory systems are

120/80mmHg and 25/10mmHg respectively.

Pulse Pressure

Arterial pulse pressure (PP) is defined as the difference between the systolic and

diastolic pressures (Equation [ 3-2 ]).

Pulse pressure is affected by a number of parameters, particularly

• Vascular resistance – Which influences runoff of blood from arteries to veins,

• Stroke volume / Heart rate – Where a reduction lowers Pdias, and;

• Compliance – This affects changes in arterial pressure with stroke volume.

Arteries are relatively compliant, thus dampening the pulsatile output of the

ventricles. Typical values of pulse pressure are 40mmHg and 15mmHg for left and

right systems respectively.

Mean Arterial Pressure

Mean arterial pressure (MAP) is related to cardiac output (CO) and systemic vascular

resistance (SVR) by Equation [ 3-3 ].

MAP is also determined by the measurement of systolic and diastolic pressures.

Since the pressure pulse wave is asymmetrical (systolic and diastolic periods

unequal), the MAP is calculated as per Equation [ 3-4 ].

The mean arterial pressures are 93mmHg for left and 15mmHg for right systems.

diassys PPPP −= [ 3-2 ]

SVRCOMAP ×= [ 3-3 ]

( )diassysdias PPPMAP −+=31

[ 3-4 ]


3-8

Venous Blood Pressure

Central venous pressure (CVP) is the pressure in the thoracic vena cava near the right

atrium. CVP, affected by changes in venous blood volume and compliance, is usually

at or near 0mmHg (Klabunde 2004). Pulmonic venous pressure (PVP) is higher than

left atrial pressure to induce flow from the lungs into to the heart (Milnor 1972).

RightAtrium

LeftVentricle

LeftAtrium

3 - 6

8-10

120

0 - 8 25

0 - 3

25

10

120

80

AORTA

RightVentricle

PulmonaryArtery

LeftVentricle

120

0 - 8RightVentricle

25

0 - 3

LeftAtrium8-10

Figure 3-4 – Pressure Distribution within the Chambers of the Heart (Rest)

Figure 3-5 – Left Heart Pressure Distribution


3-9

Mean Circulatory Pressure

Mean circulatory pressure (Pmc) is the blood pressure maintained within the

cardiovascular system when cardiac output (CO) is zero. It is directly related to the

lumped vascular compliance and blood volume. Pmc directly affects venous return

and consequently, cardiac output. For example, if blood volume is decreased,

cardiac output will reduce, causing autoregulatory mechanisms to decrease

compliance to restore Pmc until the lost volume is replaced (Guyton 1971). The

typical value of Pmc is 7mmHg (Guyton 1971; Berne, Levy et al. 2004; Klabunde

2004) and does not vary more than +/- 1mmHg for normal resting heart function

(Guyton 1971). Pmc increases with exercise due to muscle contraction, also known as

the “skeletal muscle pump”. Heart failure also results in elevated Pmc, since

autoregulation mechanisms increase blood volume and decrease venous compliance.

Chambers of the Heart

Pressures within the heart chambers vary throughout each cardiac cycle. These

pressures are described for systolic (Top) and diastolic (Bottom) phases in Figure

3-4 for resting conditions (Klabunde 2004). The pressure variations within the

chambers of the heart are described in Figure 3-3 and in further detail for the left

heart in Figure 3-5. It can be seen from Figure 3-5 that atrial pressure raises slightly

during the atrial systole phase (a), as the atria contract. Atrial pressure also slightly

increases just after ventricular systole (c) as a result of the transferral of pressure

across the closed, but compliant, mitral valve. Atrial pressure then increases slowly

during the systolic phase, as venous return increases the blood volume in the atria

(v). Ventricular pressure increases rapidly during ventricular systole, reaching a

maximum before returning to the initial diastolic phase value. The ventricular

pressure then increases to the value of atrial pressure over the remainder of the

diastolic phase, with the final pressure signifying ventricular preload.


3-10

Summary of Pressure Distribution

Figure 3-6 and Figure 3-7 graphically display the pressure distribution throughout

the systemic and pulmonic vascular trees respectively. Table 3-2 summarises the

values of pressure throughout the cardiovascular system for the condition of rest.

These values were targeted for reproduction by the mock circulation rig. Table 3-3

summarises the mean circulatory pressure values selected for resting, left heart

failure and exercise conditions in the mock circulation rig.

Figure 3-6 – Systemic Pressure Distribution Figure 3-7 – Pulmonic Pressure Distribution

Table 3-2 – Resting Pressure Distribution throughout the Vascular Tree

Table 3-3 –Mean Circulatory Pressures for Rest/LHF/Exercise

Pressures Pdiastolic (mmHg)

Psystolic (mmHg)

Mean P (mmHg)

Left Atrium (LAP) 8 10 9 Left Ventricle (LVP) 0-8 120 -

Aortic (AoP) 80 120 93 Systemic Veins (SVP) 10 10 10 Right Atrium (RAP) 3 6 4

Right Ventricle (RVP) 0-6 25 - Pulmonary Artery (PAP) 10 25 15 Pulmonary Veins (PVP) 9 12 10

Rest LHF Exercise Mean Circulatory (Pmc) 7 20 20


3-11

3.2.6 Compliance

Compliance (capacitance) is related to the ability for a vessel to distend when

encountering a change in blood volume. Figure 3-8 displays the behaviour of non-

compliant and compliant vessels in response to systolic and diastolic conditions.

Compliance is defined as a change in volume for a given change in pressure as

described by Equation [ 3-5 ] and shown in Figure 3-9. It is therefore related to the

ease with which a change in volume causes a change in pressure (Klabunde 2004).

When applied to the circulatory system, it refers to the ability for the vessels to

expand and contract under an applied pulse pressure (Figure 3-8). This feature aids

the heart in its pumping duty (Berne, Levy et al. 2004).

Arterial pressure is increased during systole and is partially maintained during

diastole due to the rebounding nature of the expanded arterial walls. The degree to

which the arterial walls rebound depends on the elastic properties of the artery, and is

deemed the arterial compliance. If the artery walls were rigid tubes, simulating low

compliance, no flow would occur during diastole. Additionally, the pressure in the

vessels would vary from an increased max systolic pressure to zero, as the

“damping” effect is removed, resulting in a greater amount of work required by the

heart. Instead, a level of compliance is observed, resulting in a continuous flow rate

of blood further down the circulatory system tree, and reduced heart work near to

that of a continuous pumping system (Berne, Levy et al. 2004).

DiastoleNon-Compliant

Systole DiastoleCompliant

Systole

Volume

Slope = Compliance

Figure 3-8 – Compliance (systole and diastole) Figure 3-9 – Definition of Compliance

PVC∆∆

=

[ 3-5 ]


3-12

Values of compliance are hard to determine, with the most common technique

relying on the measurement of pulse pressure and stroke volume. Compliance varies

throughout the population, with older people characterised with lower compliance

due to stiffening of arterial walls. Compliance can also be regulated in an individual

by altering the vascular tone, since the activation of smooth muscle fibres within the

arteries changes compliance. Compliance is not linear in biological systems, and

decreases with increasing volume, as the blood vessels have an elastic limit.

Arterial Compliance

Typical arterial compliances have been documented between 1 and 2 mL/mmHg for

the systemic arteries and 1 - 8 mL/mmHg for the pulmonary arterial system.

Therefore, these values were used as boundaries for compliance values used in the

mock circulation rig.

Arterial compliance may also be calculated theoretically if stroke volume (SV), pulse

pressure (PP) and degree of microcirculation runoff is assumed. Rapid ejection of

blood from the ventricle as it contracts expels 80% (56mL) of the stroke volume

(70mL) and takes up to 16% of the cardiac cycle (Berne, Levy et al. 2004). Since the

flow through the capillaries is constant, 16% (11.2mL) of the blood exits the arterial

side during this time. Equation [ 3-6 ] describes a calculation of systemic and

pulmonary arterial compliance, using the assumed value of PP (40mmHg).

Table 3-4 – Arterial Compliance

Systemic Arterial Compliance Pulmonary Arterial Compliance 1.14 ml/mmHg (Papaioannou, Mathioulakis et al.

2003) 1ml / mmHg (Donovan 1975)

1.0ml/mmHg (Donovan 1975) 4.8 ml/mmHg (Ferrari, Gorczynska et al. 2001) 1.44ml/mmHg (William, Stevens et al. 2003) 2-4 ml/mmHg (Milnor 1972)

1.8ml/mmHg (Ferrari, Gorczynska et al. 2001) 2 ml/mmHg (Noordergraaf 1978) 1-2 ml/mmHg (Noordergraaf 1978)

mmHgml

mmHgmlPulmonary

mmHgml

mmHgmlSystemic

98.215

8.44 Compliance rterial

12.140

8.44 Compliance rterial

==

==

A

A

[ 3-6 ]


3-13

Venous Compliance

Values of venous compliance published in the literature are given in Table 3-5.

Again, these values were used as a guide for use in the mock circulation rig.

Systemic venous compliance (SVC) is assumed to be 19-20 times that of systemic

arterial compliance (AoC) (Tyberg 2002; Berne, Levy et al. 2004; Klabunde 2004),

therefore SVC may be calculated using Equation[ 3-7 ].

Table 3-5 – Venous Compliance

Systemic Venous Compliance Pulmonary Venous Compliance 10ml/mmHg (Donovan 1975) 5ml / mmHg (Donovan 1975)

50-200 ml/mmHg (Noordergraaf 1978) (William, Stevens et al. 2003) 6 ml/mmHg (Milnor 1972)

4 ml/mmHg (Noordergraaf 1978)

mmHgml

mmHgml 4.222012.1 Compliance Venous ystemic =×=S

[ 3-7 ]

3.2.7 Resistance

The resistance to blood flow in the vascular network is predominantly derived by

vessel diameters and the network organisation (series/parallel). Changes in local

vessel diameters allow organs to regulate the amount of blood flow to meet

metabolic needs of the tissue. Vessel diameters are controlled by the secretion of

vasoconstriction and vasodilation substances (Klabunde 2004). Resistance is the

predominant factor controlling afterload, and is defined by Equation [ 3-8 ].

The majority of resistance is developed by the inherent resistance caused by the area

of the blood vessels. That is, for a particular pressure differential (P1-P2), a maximum

fluid velocity (v2) may be calculated using Bernoulli’s equation (Equation [ 3-9 ]).

QPPR 21 −= [ 3-8 ]

2

22

2

21

21

1

1

22z

gv

gPz

gv

gP

++=++ρρ

[ 3-9 ]


3-14

Since flow (Q) equals this set velocity (v2) multiplied by the vessel cross-sectional

area (A) (Equation [ 3-10 ]), altering the area (A) will result in various flow rates

(Q) for the same pressure differential, and thus resistance (R) will vary.

Total resistance should be evaluated by adding effects of factors such as fluid

viscosity (η). The nature of the flow (turbulent/laminar) will dictate the magnitude of

these effects. For example, resistance of laminar flow through a pipe vessel radius (r)

and length (l) is described by Equation [ 3-11 ] (Berne, Levy et al. 2004):

A number of different conventions are used to display values of resistance. The most

common are Peripheral Resistance Units (PRU) (mmHg.s.ml-1) and the more

medically traditional dyne.s.cm-5.

Systemic Vascular Resistance

Systemic vascular resistance (SVR), or total peripheral resistance (TPR), is a

measure of the resistance to blood flow incurred by the arteries, capillaries and veins

of the systemic circulatory system. SVR consists of the sum of each individual

component resistance (arteries, capillaries, veins), and is a measure of the difference

in mean aortic (mAoP) to right atrial pressure (RAP) divided by CO.

Values of SVR for a healthy patient from rest to exercise range from approximately

1400 to 700 dyne.s.cm-5 respectively (Burton 1965; Milnor 1982). SVR is slightly

higher for patients suffering heart failure (1596-1862 dyne.s.cm-5) due to

compensatory mechanisms (Patel, Allaire et al. 2003). Equation [ 3-12 ] describes a

sample calculation of SVR while Table 3-6 displays the guiding values for

replication in the mock circulation rig.

Table 3-6 –Systemic Vascular Resistance

5..1386min.18

5)393()(

−==

−=

−=

cmsdyneLmmHg

COmRAPMAPSVR

[ 3-12 ]

AvQ ×= 2 [ 3-10 ]

4

8rlR

πη

=

[ 3-11 ]

Rest Heart Failure Exercise 1.1 PRU

(1463 dyne.s.cm-5) 1.35 PRU

(1800 dyne.s.cm-5) 0.54 PRU

(718 dyne.s.cm-5)


3-15

Pulmonary Vascular Resistance

Pulmonary Vascular Resistance (PVR) is a cumulative measure of the resistance in

the arteries, capillaries and veins of the pulmonary system. Although the vessel

diameters are similar to the systemic system, the distance the blood must travel

within the vessels is smaller. For example, the length of the Aorta is approximately

50cm, while the Pulmonary Artery is just 5cm (Milnor 1972). PVR is therefore

expected to be much less.

Values of PVR for a healthy patient during rest and exercise conditions range

between ~50 to ~200 dyne.s.cm-5 (Milnor 1972; Slife, Latham et al. 1990). Increased

resistance to flow is again found in patients suffering heart failure. A sample

calculation of PVR is shown by Equation [ 3-13 ] while Table 3-7 specifies the

values required for replication in the mock circulation rig.

Table 3-7 –Pulmonic Vascular Resistance

Individual Component Resistances

Cardiac output is a measure of the volume of blood pumped per minute of operation.

However, since blood ejection is intermittent, the actual instantaneous local blood

flow in each section varies considerably. Establishing the actual blood flow

throughout each local section for each pulse is essential to ascertain the individual

local resistances for a given pressure. This is important when selecting the

appropriate pipe sizes throughout the mock circulation rig.

Rest Heart Failure Exercise 0.08 PRU

(106 dyne.s.cm-5) 0.12 PRU

(160 dyne.s.cm-5) 0.031 PRU

(41 dyne.s.cm-5)

5..93min.2.1

5)915()(

−==

−=

−=

cmsdyneLmmHg

COmLAPmPAPPVR

[ 3-13 ]


3-16

3.2.8 Cardiac Output

Cardiac output (CO) is dependant on a number of integrating factors. CO relates

directly to Heart rate (HR) and Stroke Volume (SV) by the following relationship

(Equation [ 3-14 ]).

Furthermore, Stoke Volume (and thus indirectly CO) is directly dependant on

• Venous Return (VR),

• Afterload, and

• Contractility

In short, an increase in VR increases end diastolic volume (EDV), a reduction in

afterload reduces end systolic volume (ESV), and an increase in contractility allows

for more blood ejection, again reducing ESV. These volume changes affect SV as

defined by Equation [ 3-15] and shown in Figure 3-10.

100

0

200

0 100

LV Volume (mL)

VenousReturn

200

ControlLoop

Figure 3-10 – Development of Ventricular Stroke Volume

HRSVCO ×= [ 3-14 ]

ESVEDVSV −= [ 3-15 ]


3-17

The reversal of these conditions (rest, high blood pressure and heart failure) results in

a reduced SV.

The most important factor affecting CO is the mean circulatory pressure (Pmc). This

pressure dictates the rate of ventricular filling during the diastolic phase, and

consequently EDV and finally SV. In other words, a low Pmc will cause reduced CO,

whereas high Pmc will increase CO considerably in a normally functioning heart.

Furthermore, an increased HR does not change CO for a given Pmc. Instead, a lower

SV occurs, since the rate of diastolic filling is unchanged, and the ventricle does not

refill to the previous ESV.

Since the heart has a limited EDV due to size and capacity, a further increase in Pmc

would cause the ventricle to fill completely early in the diastolic phase, and further

ventricular filling would cease during the latter part of the diastolic phase. To

counter this situation, HR increases, providing significant increases to CO. This is

experienced in exercise conditions, where an increased HR and Pmc are observed.

The body therefore has methods to regulate cardiac output in response to varying

physiological needs. Interactions of these techniques will be explained in further

detail in subsequent sections.

Typical values for parameters directly affecting CO under conditions of rest, exercise

and heart failure are given in Table 3-8. The table also presents the values of CO

chosen for simulation in the mock circulation rig.

Table 3-8 – Parameter affecting cardiac output

Cardiac output is often related to patient size or body surface area (BSA), and is

termed Cardiac Index (CI) (Equation [ 3-16 ]);

Typical resting values for CI range from 2.5 to 4.0 (L/min) / (m2).

Rest Exercise Left Heart Failure EDV 50 ml 50 ml 192 ml ESV 133 ml 133 ml 150 ml SV 83 ml 83 ml 42 ml HR 60 bpm 120 bpm 60 bpm

CO 4.9 L/min 9.9 L/min 2.5 L/min

BSACOCI =

[ 3-16 ]


3-18

Effect of Venous Return on Cardiac Output

Mean Circulatory Pressure (Pmc) immediately increases after the onset of exercise

due to the effect of contracting muscles, thus setting up a larger pressure gradient

back to the heart, and venous return increases accordingly. The one-way valves

inside the veins prevent retrograde flow and promote the return of blood to the heart

(Klabunde 2004). The rise in Pmc and consequent rise in SVP is halted by a change in

cardiac function. SVP is therefore maintained by a balance between heart function

and peripheral circuit conditions. The heart will adjust to an increase in venous return

during exercise via two mechanisms (Notarius and Madger 1996).

The first relates to the mechanism described by the Frank-Starling Law1, by which

the increase in venous return results in an increase in the end diastolic volume

(preload) in the ventricle, thus leading to a larger stroke volume and ultimately

cardiac output. An increase in venous return implies that a change in peripheral

circuit factors must have occurred (pressure gradient, resistance, compliance etc)

(Klabunde 2004).

The second is associated with an increase in cardiac function (contractility, heart

rate), thus leading to an increase in CO and VR. Both mechanisms can occur

simultaneously to adapt to a change in VR, however if CO is predominately due to

changes in the peripheral circuit, then the SVP will increase. Alternatively, if the CO

increase is predominantly due to changes in heart function, SVP will drop.

To balance the Pmc, both mechanisms must complement each other (Notarius and

Madger 1996). Thus an initial increase in CO is most likely due to the Frank-Starling

mechanism. A further rise in SVP is prevented by the adaptations in cardiac

response (which act to reduce SVP) until a balance is achieved as exercise continues.

________________________________

1 Frank Starling Law describes the ability of the heart to increase its force of

contraction in response to an increase in preload (cause by an increased EDV). The

increased preload stretches the cardiac myocytes allowing them to contract more

forcefully and thus eject the additional volume.


3-19

Effect of Contractility (Inotropy) on Cardiac Output

Changes in stroke volume can be achieved by the alteration in cardiac muscle

contractility. For example, improved contractility increases the rate at which

pressure is developed in the ventricle, and consequently results in additional blood

ejection. This leads to a reduced ESV, increased SV and hence increases CO.

Changes in contractility produce significant variations in ejection fraction, which is a

clinical term describing the ratio of Stroke Volume (SV) to the End Diastolic

Volume (EDV), and is used to evaluate the inotropic state of the heart. Heart failure

patients experience a reduction in contractility, leading to a reduced ejection fraction

and stroke volume. This results in a larger amount of blood remaining in the ventricle

at the beginning of diastolic filling (preload). An increase in this preload to a

pressure exceeding 20mmHg in the left ventricle can result in pulmonary congestion

and oedema, as the pressure gradient from right to left heart is reduced and

consequently flow congests in the lungs.

3.2.9 Cardiac / Vascular Coupling

Systemic Vascular Function

The heart can be described as a pump that transfers blood from a highly compliant,

low pressure reservoir (venous) to a low compliant, high pressure reservoir (arterial)

(Tyberg 2002). As previously stated, when there is zero cardiac output, the mean

circulatory pressure is approx 7mmHg (Berne, Levy et al. 2004; Klabunde 2004).

As the heart starts beating and cardiac output increases, blood is transferred from the

venous to arterial vessels. Consequently, the SVP and RAP decrease, and the aortic

pressure (AoP) increases. This produces a pressure gradient across the body vascular

resistance, causing a flow of blood from the arteries to veins. The degree that the

arterial pressure increases and RAP decrease about the Pmc is determined by the

venous (Cv) and arterial (Ca) compliance (Tyberg 2002). Again, the typical ratio of

venous to arterial compliance is nineteen (Tyberg 2002; Berne, Levy et al. 2004;

Klabunde 2004).

To illustrate this process, when CO is 5l/min, arterial pressure increases to

102mmHg (7 mmHg + 5*19) and venous pressure decreases to 2mmHg (7 - 5*1).

However, if flow was to increase further to 7 L/min, venous pressure would drop to


3-20

0mmHg and no increase in cardiac output would be possible since the resulting

pressure would be negative and the vessels would collapse (Tyberg 2002; Klabunde

2004).

Instead, the body has two alternative mechanisms to allow for greater cardiac output

and prevent the collapse of the veins. The first being an increase in total blood

volume contained in the circulation, as achieved by the conservation of salt and

water by the kidney. The second is a reduction in vascular volume by

vasoconstriction and reduction in venous compliance. Both of these regulatory

mechanisms raise Pmc. This is illustrated in Figure 3-11 where it is observed that CO

is dependant on Pmc and RAP. Furthermore, increasing SVR for a given Pmc affects

the slope of RAP to CO. This demonstrates the ability to regulate CO by changes in

vascular resistance. (Guyton 1971; Klabunde 2004).

Pmc

V Cv

V Cv

PRA (mmHg)

5

0Pmc

SVR

PRA (mmHg)

5

0

SVR

Figure 3-11 – Effect of Volume, Vascular Compliance and Resistance on PRA

Cardiac Function

According to the Frank-Starling Law, an increase in RAP increases right ventricular

preload which in turn increases CO from the right ventricle. This transfers blood

volume to the pulmonary circuit which increases LAP, and therefore increase left

ventricular preload. Consequently, Left ventricular output increases. This

autoregulatory mechanism continues until a balance is reached. As previously

identified, this process is attributed to the increase in myocyte contractility in

response to increased preload, and is illustrated as the cardiac function curve in

Figure 3-12. The path of the cardiac function curve is dependant on the inotropic

state of the heart (Guyton 1971; Klabunde 2004).


3-21

5

0

10

0 10

NormalEnhanced

Depressed

PRA (mmHg)

5

0

10

0 10

Cardiac Stimulation

PRA (mmHg)

B

C

A Cv

�

Figure 3-12 – Cardiac Function Curves Figure 3-13 – Interaction between Cardiac

and System Function Curves

Interaction of Cardiac and Systemic Function

By superimposing the systemic vascular function curve on the cardiac function

curve, the operating point of the circulation is identified (Figure 3-13). Cardiac

output is therefore altered by changing systemic and/or cardiac function. For

example, an increase in cardiac function (inotropy/heart rate) from (A) moves the

cardiac function curve up and to the left, resulting in a small increase in CO and

slight reduction in RAP (B). If however, this occurs in conjunction with a change in

systemic vascular function (e.g. reduced compliance), cardiac output is dramatically

increased (C) (Guyton 1971; Klabunde 2004). It is this ability to regulate cardiac

output by altering cardiac or systemic parameters that allows the body to

accommodate varying degrees of cardiac or systemic failure.

Figure 3-14 – Compensatory mechanisms in response to heart failure

Heart Failure

A failing heart loses its ability to pump forcefully, which translates to a loss of

contractility. Figure 3-14 describes the cardiac and systemic response to this

condition. The depressed inotropic state results in the cardiac function curve shifting

down and to the right, and consequently CO reduces to critical levels (B). This is

5

0

10

0 10PRA (mmHg)

BCA

CardiacFailure

20

V CvSVR


3-22

somewhat compensated by the Frank-Starling mechanism, as RAP is increased as a

result. However, compensatory mechanisms are also invoked by the systemic

vasculature to further increase CO. In brief, blood volume increases and venous

compliance decreases (to increase Pmc) while SVR increases to reduce the slope.

Thus, cardiac output is increased at the expense of elevated RAP (C). This is not the

ideal solution since increased RAP or LAP causes edema, as blood backs up in the

systemic (right heart failure) or pulmonary (left heart failure) systems respectively

(Guyton 1971; Klabunde 2004).

Left Heart Failure

Left heart failure leads to increased mean pulmonary pressure and reduced mean

systemic pressure. Basically, blood is transferred from the systemic to pulmonary

systems due to the healthier right ventricle. The failing left ventricle cannot

effectively remove this blood, which then “backs up” in the lungs. This excess fluid

increases the mean pulmonary pressure, and thus LAP, causing an increase in left

ventricular output due to the Frank-Starling Law. However, if this pressure increases

above 28mmHg (the colloid osmotic pressure of the plasma), fluid filters into the

interstitial spaces and alveoli. Pulmonary congestion occurs, and if remained

unchecked, a fatal condition known as pulmonary edema develops.


3-23

Right Heart Failure

The consequence of unilateral right heart failure is not severe systemic edema, but

predominantly reduced CO. Given that the pulmonary blood volume is much less

than the systemic volume, and that the compliance is much higher in the systemic

circuit, a smaller volume transfer of blood from pulmonary to systemic circuits

causes a relatively small increase in SVP. Hence, the risk of edema is low. Venous

return to the right atrium is therefore not increased and the Frank-Starling Law is not

observed. Instead, CO remains low and the potential for cardiac shock develops.

Systemic Failure

Systemic failure refers to the inability of the systemic vasculature to alter its

characteristics. This includes arteriosclerosis, which reduces the vessels compliance

and increases systemic vascular resistance.

Preload

Preload is defined as the initial stretching of ventricular cardiac myocytes prior to

contraction. As this is difficult to measure, the more common techniques used to

evaluate preload is via end diastolic volume (EDV) or end diastolic ventricular

pressure (LVPED). These methods are not ideal as they depend on the condition of

the ventricle, i.e. ventricular dilation affects EDV whereas compliance affects

pressure. Preload is dependant on the volume of blood that fills the ventricle at the

end of the passive filling diastolic heart phase. This blood volume depends on

factors such as venous pressure, venous return and heart rate (Klabunde 2004). As

already noted, the effect of increased preload is to increase CO via the Frank-Starling

Law.

Afterload

Afterload refers to the load on the ventricle caused by the vasculature, and is

predominantly dependant on the vascular resistance. The effect of afterload is to

alter the stroke volume of each heartbeat. An increase in afterload will result in

reduced blood volume ejection from the ventricle, and therefore a greater ESV. The

EDV and consequently preload then increases slightly, which stimulates the Frank-

Starling mechanism, to compensate the increased afterload by increasing heart

functionality and improving stroke volume (Klabunde 2004).


3-24

3.2.10 Autoregulation

Autoregulation of blood flow is defined as the ability of vasculature to maintain

perfusion rates despite fluctuations in pressure (Klabunde 2004). If the perfusion

pressure to an organ decreases, a corresponding decrease in flow is observed.

Metabolic or myogenic mechanisms are then activated to dilate the vessels, reducing

resistance and therefore increasing perfusion.

Autoregulation is most efficient in critical organs (Brain, myocardium). Therefore,

despite vessel constriction in response to body trauma (shock etc.), perfusion to these

critical organs is maintained. Perfusion to non-essential organs (skeletal) is reduced,

ensuring adequate oxygen delivery to critical areas (Klabunde 2004).

Regulation of cardiac output

Cardiac output is regulated by many autoregulatory mechanisms.

• Left and Right cardiac balance is automatically regulated by the Frank-

Starling Mechanism.

• A change in myocardial contractility helps to alter overall CO by adjusting

the stroke volume.

• The sympathetic and parasympathetic nervous systems regulate HR in

response to changing physiological demands.

Arterial Baroreceptors

Blood pressure in the arteries is controlled by the negative feedback system, referred

to as the “Baroreceptor reflex”. Baroreceptors are small pressure sensors located in

the aortic arch and carotid arteries. These receptors respond to changes in wall strain,

and thus pressure, by firing impulses to the brain. If the MAP or PP decrease, the

firing rate of the baroreceptors decreases and vessel vasoconstriction results, leading

to partial restoration of arterial pressure (Klabunde 2004).


3-25

3.3 Literature Review – Prototype VAD Testing

Ventricular assist programs often employ mock circulation loops to test device

iterations prior to animal and clinical trials. Present mock circulatory loops can be

generally classified as simple and non-pulsatile in nature, or advanced, including an

artificial ventricle to create pulsatile flow. This section provides a review of the

literature detailing the current techniques used to simulate the human circulatory

system in-vitro. The difficulties experienced while attempting to reproduce heart

failure in an animal model are also included, to highlight problems which may be

alleviated with a suitable mock circulation system..

3.3.1 Basic Mock Circulation Loops

The most basic loop facilitates the ability to test pumps’ performance characteristics.

I.e. pressure head, flow rate and efficiency at various rotational speeds. These loops

incorporate a flow meter, inlet and outlet pressure transducers, a resistance valve and

a reservoir (Makinouchi, Ohara et al. 1994; Yoshino, Uemura et al. 2001;

Yoshizawa, Sato et al. 2002) (Figure 3-15). Water is used as the pumping medium

at room temperature, piped through 3/8” (9.5mm) tubing (Takami, Yamane et al.

1997; Linneweber, Chow et al. 2001). The basic loop is the quickest and most

efficient means of testing a newly designed pumps’ ability to meet the physiological

requirement of the body. Although refinement of the loop by introducing different

pumping mediums improves the validation of results, it is not possible to determine

the effect on pump hemodynamic performance and operation when introduced into

the intended pulsatile system, the human circulatory system.

Figure 3-15 – Basic Test Loop

Pin

Pout Flow Meter

Valve

Reservoir

Pump


3-26

3.3.2 Advanced Mock Circulation Loops

To investigate pump performance in a pulsatile environment, it is necessary to

introduce features into the mock circulation loop to recreate parameters of the native

circulation system. Bearnson et al. (1996), Iijima et al. (1997), Pantalos et al. (1998)

and Kikugawa et al. (2000) all created mock loops incorporating; Pulsatility,

Temperature, Atrial Pressure, Arterial / Venous compliance and Peripheral resistance

(Figure 3-16). These mock loops were developed to evaluate LVAD performance,

and therefore use a systemic mock loop only. Therefore, the effect of the device on

right heart function and pulmonary circulation can not be evaluated. Furthermore,

the techniques of pulse replication do not faithfully adhere to the Frank-Starling law.

Figure 3-16 – Advanced Mock Circulation Loop

3.3.3 Current Mock Loops

The Donovan mock circulation loop was developed in 1975 (Donovan 1975) as an

in-vitro rig to test Total Artificial Hearts. Consequently, the rig does not incorporate

an artificial ventricle, and is therefore not suitable for evaluating the performance of

assist devices. It may however be adapted to further advance the basic mock

circulation loops. As such, the Donovan loop is regularly used by research centres for

VAD testing today (Bearnson, Olsen et al. 1996; Bullister, Reich et al. 2002; Chung,

Joo Lee et al. 2003; Olegario, Yoshizawa et al. 2003). Recent mock loops have

attempted to address the limitations presented. In particular, Pantalos et al. (2004)

developed an artificial ventricle that adheres to the Frank-Starling Law, however

unfortunately it is limited to the systemic circulation system only (Figure 3-17).

Pout

Flow Meter

Peripheral Resistance Valve Aortic

Compliance

Pin

Pump

Atrial Reservoir

Latex Sac

Air

Temperature Bath (37oC)


3-27

Figure 3-17 –Mock Circulation Loop (Pantalos, Koenig et al. 2004)

Wu et al. (2004) identified the need for pulmonary and systemic loop simulation

(Figure 3-18); however the artificial ventricle induces fluid into the chamber by

negative pressure, and thus does not solely rely on passive diastolic filling.

Figure 3-18 –Mock Circulation Loop (Wu, Allaire et al. 2004)

halla


halla



3-28

Physiological Condition

Both Pantalos et al. (1998) and Kikugawa et al. (2000) believe a mock circulatory

system should have the ability to alter hemodynamic characteristics to emulate

varying degrees of heart function under several physiological conditions. In light of

this, Pantalos et al. (2004) attempted to simulate normal, failing and recovering heart

function for resting conditions in a systemic only circulation loop. Furthermore, Wu

et al. (2004) used a loop to examine the reproduction of rest, sleep and exercise

conditions with five levels of heart function.

Cardiac Functionality Simulation

Pulsatile Ventricle

Heart function is sometimes limited by the use of a sucking type diaphragm or roller

pump, which creates a negative pressure to induce fluid into the ventricular chamber.

The pulsatile effect of the natural heart is critical for cardiovascular simulation. This

is particular important when evaluating ventricular assist devices. Pulsatility may be

achieved by using a pneumatically driven ventricle consisting of a flexible membrane

or piston cylinder, and tilting disc one way valves to simulate the effect of the mitral

and aortic valves (Iijima, Inamoto et al. 1997; Vermette, Thibault et al. 1998;

Kikugawa 2000; Kawahito, Takano et al. 2001; Endo, Araki et al. 2002; Olegario,

Yoshizawa et al. 2003; Patel, Allaire et al. 2003; Pantalos, Koenig et al. 2004).

However, an important requirement of any mock ventricle is its ability to reproduce

the Frank-Starling Law (Pantalos, Koenig et al. 2004). Altering the functional state

of the heart (i.e. contractility) produces a variation in preload, afterload and cardiac

output conditions (Kikugawa 2000).

Left Ventricular Pressure

Left ventricular pressure is often controlled by pulsed air pressure driven into the

space surrounding the flexible artificial ventricle. Patel et al. (2003) and Pantalos et

al. (2004) modified physiological condition by changing the driving air volume and

impulse rate, simulating contractility and time in systole respectively.

Heart Rate, Systolic Period

Patel et al. (2003) identify 60-70 bpm as a healthy patient’s HR at rest, with a

systolic period of 45%. On the other hand, heart failure patient’s HR is generally 60-

80 bpm with a systolic period of 35%. However, investigations using the mock loop


3-29

developed by Wu et al. (2004) used a common 40% systolic period and 60 bpm HR

for rest and heart failure, while HR was increased to 95 bpm for exercise conditions.

Furthermore, tests conducted on a mock loop developed by Pantalos et al. (2004)

used a common 37% systolic period for all physiological heart conditions. In this

study, HR was varied from 80 to 120bpm for each physiological condition of rest,

heart failure and recovery.

Atrial Reservoir

The atrial reservoir allows for conservation of venous return fluid momentum despite

the intermittent nature of the atrial/ventricular valves. It is located in the loop section

feeding the simulated ventricle. It can be a chamber open to atmosphere, with the

fluid level relative to the pump centreline determining LAP (Weber, Doi et al. 2002).

Vascular Simulation

Resistance

Reproducing vascular resistance is most easily reproduced by inserting valves into

the peripheral circulation branch. For example, adjustable tube clamp, screw, ball or

gate type valves were used by Weber et al. (2002) and Wu et al. (2004). However,

this technique may be further enhanced by creating a valve dependant on the arterial

pressure, as observed by the baroreceptor response in the actual human system.

Donovan (1975) and Bearnson et al. (1996) reproduced this feedback response by

connecting an outflow occluding piece to a bellows in the arterial compliance

chambers. An increase in arterial pressure (from an increased flow into the chamber)

acts to displace the volume in the bellows, resulting in a mechanism removing the

plate occluding the flow path out of the chamber, thus reducing the resistance of the

flow path and therefore reducing the arterial pressure.

Alternatively, Mohara et al. (1998) reproduced peripheral resistance by correctly

selecting the inner diameter and length of tubing to produce the desired resistance.

For example, Tayama et al. (1997) used a length (270mm) of 6.5mm (¼”) ID tubing

in a loop of 9.5mm (3/8”) ID tubing to recreate peripheral resistance equal to

100mmHg afterload. This technique replaces the resistance valve, thus reducing its

effect on blood damage. However, the ability to change resistance values is difficult.

Patel et al. (2003) identify the systemic peripheral vascular resistance of a healthy

patient as 0.85 PRU (1130 dyne.s.cm-5) at rest, while heart failure patient’s


3-30

experience a larger systemic resistance at 1.2-1.4 PRU (1596-1862 dyne.s.cm-5).

Pantalos et al. (2004) agreed that patients with heart failure encounter elevated

vascular resistance (2023 dyne.s.cm-5) from normal values (1218dyne.s.cm-5).

Compliance

Compliance of the aorta, pulmonary artery and consequent blood vessels is essential

for modelling the dynamic system for which a pump may be inserted. This enables

the system to incorporate and vary the effect of arterial elasticity and rigidity.

Donovan (1975), Bearnson et al. (1996), and Weber et al. (2002) each modelled

compliance by incorporating a closed (to atmosphere) outflow reservoir with

appropriate and variable fluid levels / air volume over the fluid. Pantalos et al.

(1998) also recommended the inclusion of a venous compliance chamber to replicate

venous compliance. The compliance chamber dimensions used by Patel (2003) were

two 1.5” (40mm) ID pipes at 16” (400mm) high, each containing a certain height of

water and air volume suitable for arterial and venous compliance. Values of

systemic arterial compliance employed by Pantalos et al. (2004) were 1.3 ml/mmHg

for both normal and failing heart mock loop simulation. Although they concede the

values were different to the clinical values of 1.4 and 0.3 ml/mmHg for normal and

heart failure respectively, they maintain both vascular input impedance, as well as

mechanical properties are the most critical parameters.

Cardiac Output

To simulate a failing heart, Endo et al. (2002) suggests cardiac output should vary

from 0.5 - 2.1 L/min. Both Patel et al. (2003) and Wu et al. (2004) recommend

values of 2 – 4.4 L/min for failing heart simulation, 5.2 L/min for normal heart

function and 7.95 L/min for exercise conditions. Pantalos et al. (2004) simulated 3

L/min for failing and 5 L/min for normal hearts at rest, while Golding and Smith

(1996) proposed that for normal heart simulation, cardiac outputs are expected

between 3.75-6.75 L/min at rest, and 5.6-10L/min under moderate activity.

Pressure Distribution

Patel et al. (2003) predict the resting mean aortic pressure (MAP) of normal patients

is 90-140mmHg, and slightly lower during heart failure. Wu et al. (2004) reaffirm

this observation in a study that attempted to reproduce five degrees of heart function.

A MAP of 102mmHg was identified as normal heart function, with congestive heart


3-31

failure simulated at 53mmHg. Additionally, a LVPED of 3-12 mmHg is considered

normal, with values above this range expected in congestive heart failure patients.

The normal heart function in the study encountered a LVPED of 9mmHg, with

18mmHg experienced in the same heart failure condition. The study by Pantalos et

al. (2004) demonstrated a normal MAP of 95mmHg while heart failure was

reproduced as 65mmHg. LVPED was also recorded at 2-5mmHg and 15-35mmHg

for normal and heart failure simulations respectively.

Cardiac Assist Device Insertion

Chung et al. (2003) summarise the guidelines set by the National Institutes of Health

(NIH), which call for the preload on the left and right ventricles of a normal heart to

be 10mmHg, while afterload is 20mmHg on the right and 100mmHg on the left

ventricle. Furthermore Weber et al. (2002) maintains the inserted pump must be

capable of at least 8 L/min output.

To determine the effect of preload on the assist device, Endo et al. (2002) and

Weber et al. (2002) recommend the level of fluid (RAP and LAP) in the atrial tank

should be varied (-6 to 30mmHg) to simulate a change in venous return to the heart.

Tests may also be conducted with various right and left afterload.

Weber et al. (2002) further isolate the effect of afterload variation (vascular

resistance) by shunting both atrial reserves to keep the pressures at an identical level

of 13-15 mmHg. Each side must then be tested independently. The afterload should

be kept constant on the non-test side while the afterload on the tested side is varied

(Right = 20-40mmHg, Left = 60-170 mmHg).

Endo et al. (2002) suggest the sensitivity of the circuit to preload, afterload and

contractility variations induced by various stages of heart failure should be

investigated to determine the potential for ventricular sucking. Yuhki et al. (1999),

Olegario et al. (2003) and Wu et al. (2003) also agree the potential for the rotary

pump to collapse the ventricle should be investigated. If the simulated ventricle is

constructed from a suitably collapsible material (silicon balloon/latex rubber),

inserting the inflow cannula into directly into the ventricle will provide evidence of

sucking, given that the compliance of the ventricle material is similar to that of the

natural heart. Alternatively, Iijima et al. (1997) indicate a small collapsible tube

between the inflow to the pump and the ventricle will also reproduce this effect.


3-32

Pumping Mediums

Water is the most common pumping medium used in circulation loops. Water is the

most efficient medium to conduct tests, as it is cheap and readily available and the

density (1000 Kg/m3) is similar to that of the intended medium, blood (1053 Kg/m3).

Pantalos et al. (1998) and Yoshino et al. (2001) each used saline solutions (0.9%

NaCl) to closely match the aggressive, corrosive nature of the physiological

environment, while glycerol solutions at 37% (Kawahito, Takano et al. 2001;

Yoshizawa, Sato et al. 2002), 40% (Ayre, Vidakovic et al. 2000), 45% (3.5 cP)

(Gerhart, Horvath et al. 2002) or 52% (Chung, Joo Lee et al. 2003) have been stated

to replicate the viscosity of blood. Ayre et al. (2000) claim a Glycerine-water

analogue is the ideal solution to vary viscosity over a large range. Properties of the

mediums used in mock loops are given in Table 3-9.

Table 3-9 – Properties of Pumping Mediums

Density (Kg/m3)

Dynamic Viscosity (mPas)

Water 998 1.0 Saline 1025 1.0

Haemaccel 3.5% (0% hematocrit) 1050 1.2

Haemaccel (39% hematocrit) 1050 3.9

Blood 1050 4 Glycerol (40%) 1103 3.9

The use of these solutions does not however provide evidence of pump performance

with regard to blood damage. Nonetheless, these mediums allow for any material to

be present in the pump. Therefore initial designs are tested with these solutions to

determine their basic performance.

Once a final design is manufactured from biocompatible materials, it is inserted into

a basic loop containing blood. To this end, Takami et al. (1997) and Yoshikawa et al.

(1999) circulated 400-450ml of animal (Bovine) or human blood. Tests of this

nature are governed by ASTM standard F1841-97, where the loop is initially filled

with a sterile saline solution and circulated for 30min then replaced with test blood

(ASTM_F1841-97 1998; Linneweber, Chow et al. 2001).

This standard specifies a basic mock circulation loop for the purpose of testing pump

hydraulic performance and blood damage level. The closed loop (Figure 3-19)

consists of two meters of 9.5mm ID PVC tubing, a 13cm x 13cm reservoir with


3-33

sampling port, a screw clamp for resistance, Ultrasonic or Magnetic Flow meters, a

Thermistor and 450ml +/- 45 ml primed blood volume. Pantalos (1998) recommend

the fluid temperature be maintained at 37oC +/- 3oC to correctly simulate

physiological conditions. Tayama (1997) achieve this by running a portion of the

tubing through a warm water bath.

Figure 3-19 – ASTM F1841Test Loop (ASTM_F1841-97 1998).

Instrumentation

Mechanical assist devices inserted into the loop must be evaluated in terms of their

performance. Tansley (2000) explain this performance relates to flow rate, pressure

head development, efficiency (system and stage), with and without a pulsatile

environment. Gerhart (2002) advise pressure measurement locations should be

within six inches of the pump ports to eliminate effects of head friction losses.

Takami (1997) and Yoshino (2001) attained these measurements a simple U-tube

manometer, while Tansley (2000) used disposable and non-disposable pressure

transducers

halla



3-34

Cardiac output is best monitored using non-fluid contacting probes. Yoshikawa et al.

(1999) and Yoshizawa et al. (2002) utilised magnetic probes Model 540, (Medtronic

Biomedicus, Inc., USA), while the more popular TRANSONIC ultrasonic flow

probes models are used by many centres (Takami, Yamane et al. 1997; Linneweber,

Chow et al. 2001; Weber, Doi et al. 2002; Chung, Joo Lee et al. 2003). These flow

probes provide the advantage of low flow resistance and non-contact measurement,

conventionally not available in paddle or optical flow probes.

3.3.4 In-Vivo Reproduction of Heart Failure

The importance of including variable left and right heart functionality in the mock

circulation loop is highlighted by the difficulties experienced in reproducing heart

failure in an in-vivo animal model.

Reitan et al. (2002) revealed that assessment of LVAD performance in-vivo is often

conducted in an animal with normal heart function. The results do not always

correlate to the evaluation of circulatory support in a heart failure model, and it is

therefore difficult to predict the hemodynamic contribution of the implanted device.

Consequently a number of techniques have been attempted to induce heart failure.

Myocardial Infarction is induced by ligation of coronary arteries; however the degree

of ventricular damage is unpredictable. Implementing pacemaker induced

tachycardia, coupled with inotropic drugs often leads to complete bi-ventricular

failure, rather than the desired left ventricular failure only. These procedures are

therefore not particularly suitable for LVAD assessment.

The group understood these difficulties and developed an adjustable left heart failure

model. Inotropic drugs were administered to induce bi-ventricular failure, and a

centrifugal pump was inserted to bypass and restore the function of the failed right

ventricle. The pump was set to output a continuous flow regardless of pulmonary

afterload. Therefore, left ventricular filling pressures / left atrial pressures were

increased from 5mmHg to 25mmHg, due to the backward congestion caused by the

failed left ventricle.


3-35

3.4 Simulation Model of Circulatory System

SIMULINK models based on underlying mathematical principles are an important

tool for evaluating a new design. By modelling a system, variables and parameters

can be easily altered to determine the magnitude of their effect on the overall system.

In terms of the mock circulation loop, it is possible to create a working model of the

system that incorporates variable vascular parameters for compliance and resistance,

together with heart functionality in the form of contractility and heart rate. This

enables greater insight into the construction and operation of an experimental test rig.

Therefore, parameter values were taken from the physical mock circulation rig to

assist design and provide a mechanism for validation.

This section details the development and operation of the complete cardiovascular

system model, before results are presented for simulated degrees of heart function

under specified physiological conditions.

3.4.1 Simulation Circulation System Configurations

The complete circulatory system loop was modelled with conditions of rest and heart

failure imposed. Individual components (heart, arteries, and veins) of both systemic

and pulmonary circuits were connected in series (Figure 3-20).

Complete Circulatory System

VENACAVA

BODY

RIGHT HEART

PULMONARY VEIN

PULMONARY ARTERY

LUNGS

LEFT HEART

AORTA

Analysis

Figure 3-20 – SIMULINK Model for Complete Circulation


3-36

30 31 32 33 346

8

10

Time (sec)

Pre

ssur

e (m

mH

g)

(b) Left Atrial Pressure

LAP

30 31 32 33 340

50

100

Time (sec)

Pre

ssur

e (m

mH

g)

(c) Left Ventricular Pressure

LVP

30 31 32 33 34

80

100

120

Time (sec)

Pre

ssur

e (m

mH

g)(d) Aortic Pressure

AoPMAP

30 30.2 30.4 30.6 30.80

20

40

60

80

100

120

Time (sec)

Pre

ssur

e (m

mH

g)(a) Systemic Pressure Distribution

LAPLVPAoPMAPSVP

30 31 32 33 34

222324

Time (sec)

Pre

ssur

e (m

mH

g)


LAP

30 31 32 33 340

50

100

Time (sec)

Pre

ssur

e (m

mH

g)


LVP

30 31 32 33 3440

60

80

Time (sec)

Pre

ssur

e (m

mH

g)

(d) Aortic Pressure

AoPMAP

30 30.2 30.4 30.6 30.80

20

40

60

80

100

120

Time (sec)

Pre

ssur

e (m

mH

g)

(a) Systemic Pressure Distribution

LAPLVPAoPMAPSVP

(a) Systemic Pressure Distribution (a) Systemic Pressure Distribution

30 31 32 33 34

2468

Time (sec)

Pre

ssur

e (m

mH

g)

(b) Right Atrial Pressure

RAP

30 31 32 33 340

20

40

Time (sec)

Pre

ssur

e (m

mH

g)

(c) Right Ventricular Pressure

RVP

30 31 32 33 3410

20

30

Time (sec)

Pre

ssur

e (m

mH

g)

(d) Pulmonary Artery Pressure

PAPMPAP

30 30.2 30.4 30.6 30.80

5

10

15

20

25

30

35

40

Time (sec)

Pre

ssur

e (m

mH

g)

(a) Pulmonary Pressure Distribution

RAPRVPPAPMPAPPVP

30 31 32 33 3415

20

Time (sec)

Pre

ssur

e (m

mH

g)


RAP

30 31 32 33 340

20

40

Time (sec)

Pre

ssur

e (m

mH

g)


RVP

30 31 32 33 34

30

40

Time (sec)

Pre

ssur

e (m

mH

g)


PAPMPAP

30 30.2 30.4 30.6 30.80

5

10

15

20

25

30

35

40

Time (sec)

Pre

ssur

e (m

mH

g)

(a) Pulmonary Pressure Distribution

RAPRVPPAPMPAPPVP

(b) Pulmonary Pressure Distribution (b) Pulmonary Pressure Distribution

Figure 3-21 – Resting Pressure Distribution Figure 3-22 – Left Heart Failure Pressure Distribution

30 31 32 33 340

10002000

Time (sec)Flow

Rat

e (m

L/se

c)

(b) Left Ventricular Filling

LAQ

30 31 32 33 340

5001000

Time (sec)Flow

Rat

e (m

L/se

c)

(c) Left Ventricular Ejection

LVQ

30 31 32 33 34456

Time (sec)Flow

Rat

e (L

/min

)

(d) Systemic Capillary Flow

SQMSQ

30 31 32 33 340

100200

Time (sec)Flow

Rat

e (m

L/se

c)

(e) Systemic Venous Return

SVQ

30 30.2 30.4 30.6 30.80

200

400

600

800

1000

1200

1400

Time (sec)

Flow

Rat

e (m

L/se

c)

(a) Systemic Flow Distribution

LAQLVQSQMSQSVQ

30 31 32 33 340

5001000

Time (sec)Flow

Rat

e (m

L/se

c)

(b) Left Ventricular Filling

LAQ

30 31 32 33 340

200

Time (sec)Flow

Rat

e (m

L/se

c)

(c) Left Ventricular Ejection

LVQ

30 31 32 33 34234

Time (sec)Flow

Rat

e (L

/min

)

(d) Systemic Capillary Flow

SQMSQ

30 31 32 33 340

100200

Time (sec)Flow

Rat

e (m

L/se

c)

(e) Systemic Venous Return

SVQ

30 30.2 30.4 30.6 30.80

200

400

600

800

1000

1200

1400

Time (sec)

Flow

Rat

e (m

L/se

c)

(a) Systemic Flow Distribution

LAQLVQSQMSQSVQ

(a) Systemic Flow Distribution (a) Systemic Flow Distribution

30 31 32 33 340

500

Time (sec)Flow

Rat

e (m

L/se

c)

(b) Right Ventricular Filling

RAQ

30 31 32 33 340

500

Time (sec)Flow

Rat

e (m

L/se

c)

(c) Right Ventricular Ejection

RVQ

30 31 32 33 342468

10

Time (sec)Flow

Rat

e (L

/min

)

(d) Pulmonary Capillary Flow

PCQMPCQ

30 31 32 33 340

500

Time (sec)Flow

Rat

e (m

L/se

c)

(e) Pulmonary Venous Return

PVQ

30 30.2 30.4 30.6 30.80

100

200

300

400

500

600

Time (sec)

Flow

Rat

e (m

L/se

c)

(a) Pulmonary Flow Distribution

RAQRVQPAQMPPCQPVQ

30 31 32 33 340

500

Time (sec)Flow

Rat

e (m

L/se

c)

(b) Right Ventricular Filling

RAQ

30 31 32 33 340

500

Time (sec)Flow

Rat

e (m

L/se

c)

(c) Right Ventricular Ejection

RVQ

30 31 32 33 3412345

Time (sec)Flow

Rat

e (L

/min

)

(d) Pulmonary Capillary Flow

PCQMPCQ

30 31 32 33 340

200400

Time (sec)Flow

Rat

e (m

L/se

c)

(e) Pulmonary Venous Return

PVQ

30 30.2 30.4 30.6 30.80

100

200

300

400

500

600

Time (sec)

Flow

Rat

e (m

L/se

c)

(a) Pulmonary Flow Distribution

RAQRVQPAQMPPCQPVQ

(b) Pulmonary Flow Distribution (b) Pulmonary Flow Distribution

Figure 3-23 – Resting Perfusion Figure 3-24 – Left Heart Failure perfusion


3-37

A demo SIMULINK model ‘PYSBE’ obtained from MATLAB was sourced and

extensively refined to mimic the configurations of the experimental circulation rig.

Refinements included:

• Elimination of temperature variation effects

• Reduction / combination of peripheral resistances into lumped resistances

• Alteration of Ventricular contraction simulation

• Addition of atrial subsystems

• Modification of all component parameters to independently researched

clinical values and corresponding actual test rig values

• Introduction of a variable mean circulatory pressure Pmc.

A comprehensive description and breakdown of the SIMULINK model is provided

in Appendix D.

3.4.2 Results

Results were obtained for pressure and flow rate distribution throughout the

simulated complete circulation system under conditions of rest and heart failure.

The pressure distribution throughout the various components of the systemic

vascular tree reproduced by the complete SIMULINK circulation loop is shown in

(Figure 3-21). A mean aortic (MAP) pressure of 97mmHg and pulmonary arterial

(PAP) pressure of 17mmHg was simulated. Furthermore, LAP varied from 7-

10mmHg, while RAP oscillated about 3mmHg.

Simulated left heart failure encountered an expected decline in MAP to 59mmHg,

and rise in PAP to 30mmHg (Figure 3-22). LAP was also predictably higher

(23mmHg).

Systemic and pulmonic flow rates were recorded during simulated resting (Figure

3-23) and left heart failure (Figure 3-24) conditions. Perfusion was reduced from a

normal value (5.1 L/min) to those expected during LHF (2.67 L/min). Individual

component flow rates were also investigated over a full cardiac cycle. The

maximum ventricular ejection flow rate for simulated rest was 400 and 310 ml/sec

for left and right ventricles respectively. This was reduced to 200 and 190 ml/sec

when left heart failure was replicated.


3-38

Figure 3-25 – Mock Circulation Rig


3-39

3.5 Experimental Mock Circulation Loop

Many current mock circulation systems concentrate solely on left ventricular assist

device evaluation, and may not incorporate pulsatile flow. This section describes the

development of a complete mock circulation loop, which incorporates pulsatile left

(systemic) and right (pulmonic) circulatory systems (Figure 3-25), designed to

function under the conditions of rest, exercise and heart failure. The equipment and

techniques used to reproduce design criteria of heart functionality, vasculature

performance, hemodynamic conditions and physiological conditions are detailed.

Experimental methods for measuring such performance characteristics are then

described in detail. Finally, the validity of the circulatory system with respect to

clinical hemodynamics is attained by presenting performance results during the

various simulated physiological conditions.

3.5.1 Design Criteria

To investigate the effect of Left- Right- and Bi- Ventricular assistance on the human

cardiovascular system, a systemic and pulmonary mock loop that adheres to the

Frank-Starling law is required. Additionally, the mock circulatory system should

have the ability to alter its hemodynamic characteristics to emulate varying degrees

of left and/or right heart function (normal, enhanced, failing) under several

physiological conditions (rest, exercise). Finally, there should be provision to easily

introduce any form of mechanical assistance.

3.5.2 Description of Mock Circulation Loop

Successful development of a circulation loop that closely simulates the native

cardiovascular system requires a comprehensive knowledge of cardiovascular

anatomy and hemodynamics. These anatomical features must be recreated in detail,

and techniques must be introduced to control hemodynamics and measure

performance.

This mock loop was developed to investigate the effect of Left- Right- and Bi-

Ventricular assistance on the human cardiovascular system. The rig follows the

Frank-Starling response to all physiological conditions by design, and features

reproducible and independently variable levels of left and/or right heart function.

Easily variable vascular parameters are included to dictate clinically expected


3-40

hemodynamic values. Most importantly however, systemic and pulmonary

circulation loops are connected in series. The circulation path and rig components are

detailed in Figure 3-26.

Figure 3-26 – Top View Mock Circulation Loop with listed parameters

01 Left Ventricle 08 Tricuspid Valve 15 Left Atrium 02 Aortic Valve 09 Right Ventricle 16 Mitral Valve 03 Systemic Arterial

Compliance 10 Pulmonary Arterial Valve 17 Right Air Compressor

04 System Flow Meter 11 Pulmonic Arterial Compliance

18 Left Air Compressor

05 Systemic Pinch Valve Resistance

12 Pulmonary Flow Meter 19 Left Solenoid

06 Systemic Veins 13 Pulmonary Pinch Valve Resistance

Vx Ball Valves

07 Right Atrium 14 Pulmonary Venous Compliance

Circulation Rig Configurations

Although the mock loop was designed to simulate the complete cardiovascular

system, the ability to independently recreate systemic or pulmonary circuits was

introduced by the actuation of strategically placed ball valves. This extends the

ability of the loop to recreate previously designed mock circulatory systems that

concentrate solely on left or right assistance. In this configuration, the system can

also function as a basic mock loop, thus allowing the evaluation of non-pulsatile

performance characteristics.

08

1011

14

12

1307

V5

V6 V7

V8

03

06

V3

04

05

02

01

16

15

V1

19

1817

V4

V2

09


3-41

Table 3-10 details the valve actuation status to reproduce the simulated left, right or

total system.

Table 3-10 – Valve Settings for Various Circulation Configurations

Total System Left Only Right Only Valve 1 ON OFF OFF Valve 2 ON OFF OFF Valve 3 OFF ON N/A Valve 4 OFF N/A ON

Cardiac

Structural

The left and right hearts consist of passive atrial and pneumatically actuated

ventricular chambers divided by check valves. The ventricles are further isolated

from the arterial vasculature by another check valve (Figure 3-27).

Figure 3-27 – Structural components of the simulated beating heart

Atrial Chambers These chambers are open to atmosphere and constructed from 40mm clear flexible

PVC piping to provide sufficient compliance (change in volume / change in water

height) and allow fluid level viewing. In short, an extra water volume corresponds to

a water level rise, which corresponds to a pressure head. Therefore the compliance is

the volume change (mL) / the change in pressure head (mmHg). A transducer is

fitted to a port at the base of the chamber to provide the atrial pressure measurement

as a head of water. A sufficient atrial pressure must be maintained to provide a

sufficient pressure gradient during the ventricular refilling phase.


3-42

Mitral / Tricuspid Valves One way swing check valves are inserted between the atrial chamber and the

following ventricle chambers. The resistance to flow of these valves must be low

enough so as to allow unimpeded and rapid flow from atrium to ventricle (required

by phase 4a – rapid ventricular filling of the cardiac cycle). Although average native

valves are approx 23mm in diameter, 40mm brass check valves were used since the

swing component has a resistive weight and obstructs some of this area.

Ventricular Chambers The ventricular chambers are similar in construction to the atrial chambers; however

they are sealed and tapped with a 6.5mm hose-tail enabling compressed air to be

input during systole and vented during diastole. A pressure transducer is fitted

together with a small compliance chamber (syringe) to reduce the effects of water

hammer experienced by the check valves.

Aortic / Pulmonic Valves Brass swing check valves were placed after each ventricle to prevent backflow from

the arterial line during the isovolumetric relaxation (diastolic) phase.

Functional

The components employed for the reproduction of heart function are displayed in

Figure 3-28.

Regulator

Pressure Gauge

3/2Solenoid

Figure 3-28 – Control and monitor of cardiac function


3-43

Heart Contraction The simulated heart must be able to pump a volume of fluid at a certain pressure for

each contraction. These pressures are reproduced by regulating (Model 11-818-100,

Norgren) compressed air and introducing it into the top of the left and right

ventricular chambers during the systolic phase. Venting the air during the diastolic

phase allows refilling of the ventricle, with the resulting water level in the chamber

representing the end diastolic volume and pre-load (head of water). Thus ventricular

pressures rise to systolic levels, before falling to zero gauge pressure at the onset of

diastole. Left and right end diastolic pressure depends on left and right atrial

pressure. However, this final value may also be influenced by the solenoid vent area.

Contractility The contractility (rate of pressure rise) is controlled by the mass flow rate of air into

the chamber, as this determines stroke volume for each contraction. This may be

varied by restricting the flow of air out of the compressor and into the regulator (in

the form of another regulator to reduce the air pressure leaving the compressor), thus

restricting the air flow into, but not out of, the ventricular chambers. “Relaxation”

(rate of pressure drop) is also important and relates to the ability for air contained in

the ventricular chamber to escape. This can be varied by using a valve to change the

effective area of the solenoid exit port.

Heart Rate and Systolic Time The rate of contractions must be variable and controllable. This is achieved by

utilising a 3/2 solenoid valve for each ventricle. Signalling the solenoid ON allows

the compressed air into the chamber simulating the systolic phase, while signalling

the solenoid OFF vents the air within the ventricle during diastole. The switching

rate determines the heart rate while the percentage time ON and OFF during each

cycle influences the systolic and diastolic times, and consequently the volume of air

input into the ventricle. The solenoid effective area when open and venting must be

sufficient to allow suitable air flow into and out of the ventricle, thus an SMC-317,

3/2 Solenoid with an effective area of 26mm2 was used.


3-44

Vasculature

Compliance

A simple loop from left ventricle to right atrium and right ventricle to left atrium

would serve the function of systemic and pulmonic arteries and veins respectively.

However, the stiffness and rigidity of the 25mm poly piping used does not compare

to the elastic nature of blood vessels, and physiological pressures of 120/80 mmHg

would not be observed. Therefore, to reproduce the compliant nature of such vessels,

arterial and venous chambers are plumbed into the circuit. Each chamber is closed to

atmosphere, and contains a certain volume of air above the water level. This volume

of air controls the change in chamber pressure with the introduction of an extra

volume of water into the chamber (compliance). Equation [3-17] characterises the

change in pressure of a vessel due to the introduction of an additional amount of air

to the initial air volume. The process is assumed isothermal for simplicity.

This can be rearranged to solve V1 by substituting for V2 (Equation [ 3-18 ]).

Since the arteries and veins of the systemic and pulmonic systems have different

pressure changes for given stoke volume (i.e. compliance values), chambers of

appropriate dimensions were used recreate the desired compliances (Table 3-11).

The initial compliance was reduced during operation since fluid is shifted to create

the diastolic value of pressure, therefore increasing the density and thus reducing the

physical volume of air within the chambers. Compliance values within the chambers

can alter with differing physiological conditions, therefore arterial and venous

chambers were capped with movable 100mm and 150mm test plugs respectively

(Figure 3-29). These test plugs had a thickness of 40mm and could be easily moved

up and down the pipe to change the initial volume of air (IVair) contained above the

water and thus effectively alter compliance.

2211 VPVP = [3-17]

P1 = Diastolic Pressure (Pdias) V1 = Initial volume of air above the fluid (IVair) P2 = Systolic Pressure (Psys) V2 = Volume of air above the fluid after systole (IVair – SV)

1

121

)(P

SVVPV −=

[ 3-18 ]


3-45

Table 3-11 – Chamber characteristics for Resting Compliance Values

Table 3-12 – Mock Circulation Rig Component Inherent Resistance Values

Systemic Arterial

Pulmonic Arterial

Systemic Venous

Pulmonic Venous

Volume Change (mL) 56 56 10 10

Pdias (mmHg) 80 10 10 9

Psys (mmHg) 120 25 10.2 12

Tank Dimensions (mm) 100φ x 350 100φ x 500 150φ x 1000 100φ x 1000

Height test plug above fluid (mm) 230 300 920 720

Initial Volume of Air (mL) 1,805 2355 17,622 5,652

Diastolic Air Volume (mL) 1,621 2325 17,490 5,588

Target Compliance (mL/ mmHg) 1-2 2-3 20-50 4-6

Initial Compliance 2.3 3 23.25 7.4

Operational Compliance 1.85 2.94 23 7.1

Components Rest (PRU)

LHF (PRU)

Exercise (PRU)

Tricuspid Valve 0.004 0.004 0.004 Pulmonary Artery Valve 0.0156 0.0156 0.0156 Pulmonary Capillaries 0.061 – Inf 0.061 – Inf 0.061 – Inf

Pulmonary Veins 0.0041 0.0041 0.0041 Mitral Valve 0.0049 0.0049 0.0049 Aortic Valve 0.0277 0.0277 0.0277

Systemic Capillaries 0.256 – Inf 0.256 – Inf 0.256 – Inf Systemic Veins 0.008 0.008 0.008

Figure 3-29 – Compliance Chamber Figure 3-30 – Proportional Pinch Valve


3-46

Resistances

The mock loop must account for the peripheral vascular resistances of the systemic

and pulmonary circulatory systems. Each arterial, capillary and venous resistance

was lumped into individual systemic (SVR) and pulmonic (PVR) vascular resistance

values respectively.

Initial SVR and PVR were experienced by the inherent area resistance of the 25mm

poly piping, check valves and fittings. This was further refined by proportional

control pinch valves (ECPV-375B, HASS Manufacturing, USA) located between the

arterial and venous chambers on their respective circulatory sides (Figure 3-30).

The valves pinch 9.5mm (3/8”) ID tubing in proportion to a 1-5 V input signal, thus

allowing the overall resistances to be finely adjusted between 700 - 1500 dyne.s.cm-5

and 77-250 dyne.s.cm-5 for systemic and pulmonary loops respectively.

Inherent resistance values were calculated by taking the required pressure drop

across a component and dividing it by the maximum flow rate through the

component (Equation [ 3-8 ]). To reiterate, maximum flow rate is calculated by first

using Bernoulli’s equation (Equation [ 3-9 ]) to determine maximum fluid velocity

(v2), assuming initial and final water heights are equal (z1=z2) and initial velocity (v1)

is zero. Multiplying final velocity (v2) by each component’s maximum pipe cross

sectional area (A) reveals maximum flow rate Equation [ 3-10 ].

Using this method, the minimum inherent resistance of the tubing due to the internal

cross sectional area was calculated at 323 dyne.s.cm-5 (0.256 PRU) for systemic, and

77 dyne.s.cm-5 (0.061 PRU) for pulmonary loops. Values calculated for each rig

component are displayed in Table 3-12. Further reductions in flow rate and therefore

increase in resistance would occur as a result of frictional resistance to flow, however

these losses are assumed to be negligible as flow is laminar, viscosity is low and pipe

section length is small.

The use of pinch valves proportionally controlled by a voltage input signal allows,

with a suitable control algorithm, the resistance to be varied in response to changes in

arterial pressure, as regulated by the baroreceptors of the natural system. This was

not incorporated in this study, and therefore forms a task for future implementation.


3-47

Hemodynamic Performance

Pressures

Exact physiological pressures must be reproduced within each simulated chamber

and blood vessel.

Table 3-2 describes the desired distribution of pressures throughout the circulatory

system for conditions of rest. Pressures were attained by the actuating ventricle and

ranges maintained by the appropriate compliance and resistances. Values of Venous

and Atrial pressure will be slightly higher for conditions of heart failure.

Mean circulatory pressure (Pmc) was achieved by pressurising the system to the

required level (Table 3-3) by adding a volume of fluid to the open atrium prior to

commencement of the heart beat.

Pressures were monitored and recorded via two methods. Firstly, qualitative

measurements were made with Edwards Life Sciences (NPC-100, GE, USA)

disposable pressure transducers coupled to a patient monitor (Series 7010, Marquette

Electronics, USA). Quantitative results were obtained using TOYODA (SD10B-1,

Gambro, USA) pressure transducers supplied with a 12V source returning a

calibrated voltage signal.

Flow Rates

Physiological perfusion rates must be simulated in the mock circulation loop

circulation. Three flow rates, determined from clinical setting and literature, are

targeted for reproduction (Table 3-13). Systemic and pulmonary flow rates were

independently monitored with two electromagnetic flow meters (IFC010, KROHNE,

Netherlands) (Figure 3-31).


3-48

Figure 3-31 – Electromagnetic Flowmeter

Table 3-13 – Physiological flow conditions

Rest Exercise Heart Failure Flow Rate 5 L/min 9 L/min 2-3 L/min

Simulated Physiological Conditions

The physiological conditions of Rest, Exercise and Heart failure are reproduced by

changing the mean circulatory pressure (Pmc), altering heart contractility (by

regulating the air pressure from the compressor), resistances (by changing the pinch

valve restriction), and arterial compliance (moving test plugs within chambers). The

aim was to reproduce clinical hemodynamic response to each condition.

Rest

Setting the vascular parameters (Pmc, compliance, resistance) as well as heart

function (Contractility and heart rate) to values for rest determined previously,

cardiac output and pressure distribution was attained for the resting condition.

Exercise

An exercising body is characterised by an increased perfusion rate as a result of

increased Pmc, reduced vascular resistance, increased MAP, and enhanced heart

function (rate and contractility). Furthermore, ventricular filling during diastole also

relies heavily on atrial contractions. Unfortunately, not all changes required for

exercise could be reproduced in the complete system.


3-49

For example, although SVR could be sufficiently lowered, minimum inherent

resistance of the pulmonary system was larger than that required of PVR for

exercise. Therefore, the exercise condition could only be represented in the systemic

loop configuration only. This concession influenced either diastolic ventricular

filling or perfusion rates. The former condition arises if the value of LAP is lowered

to RAP. Alternatively, the LAP could be maintained at the expense of reduced

systemic pressure gradient and thus the later condition is evident. Therefore a

compromise pressure was identified.

Another drawback of the current loop involves the absence of atrial kick

reproduction. However, it is entirely possible to reproduce this natural atrial function

by scaling down the technique for ventricular contractility. The atrial chambers can

be capped and compressed air (regulated to lower pressure) introduced for a period

during the diastolic cardiac phase.

Heart function was successfully replicated for exercise conditions. Heart rate was

increased by altering the switching rate of the solenoid valve, and SVR was

sufficiently reduced by increasing the aperture in the pinch valves to allow the

increased perfusion. Contractility was enhanced by increasing the value of regulated

ventricular chamber input pressure.

Heart Failure

Correct heart failure simulation depends on the degree of failure, as well as the level

of medical therapy treatment. For example, in untreated heart failure, increasing

SVR is the natural autoregulatory response to maintain MAP at the expense of

perfusion rate. However, heart failure medical treatment reduces SVR to improve

perfusion, resulting in lower MAP values.

Varying degrees of Left Heart Failure, Right Heart Failure and Bi-Ventricular Heart

Failure are possible by reducing the contractility (i.e. mass air flow rate) of one or

both of the individual ventricles, increasing Pmc (by adding more fluid volume),

reducing chamber compliances (by lowering test plugs), and adjusting peripheral

vascular resistance (by actuating the pinch valves).


3-50

Figure 3-32 – Experimental Technique


3-51

Experimental Technique

CONTROLDESK was used as an interface to a SIMULINK model that

communicated with the rig via a dSPACE card. This enabled the capture of

hemodynamic variables (such as flow and pressure) and real time control of some

vascular and heart parameters (vascular resistance, heart rate/systolic period) (Figure

3-32).

Control Software

MATLAB and SIMULINK (Mathworks) software packages were used to develop

the control model for circulation rig and mechanical assist device operation. The

complete model is extensively reviewed in Appendix D. The model allowed for the

real time control of the solenoid valves (Heart rate and Systolic period), proportional

control valves (vascular resistance), and assist device motor speed, using the

CONTROLDESK interface.

DSP Board

Digital signal processing was accomplished using a dSPACE PCI signal processing

board (DS1104, MI, U.S.A). This model accommodates 8 D/A (2 x Pinch Valve), 8

A/D (6 x Pressure, 2 x Flow) signals, 2 encoder inputs (1 x assist device motor

speed) and 2 digital I/O ports (2 x Solenoid Valve).

Instrumentation

The instrumentation used to measure rig component performance retuned voltage

signals in proportion to each measured value. These signals were displayed and

captured in real time with CONTROLDESK.


3-52

3.5.3 Experimental Procedure

A description of the experimental procedure for mock loop evaluation follows.

Parameter values are referenced in Table 3-14.

Table 3-14 – Mock Circulation Loop Parameter

Rest LHF Exercise Left Right Left Right Left

Heart Functionality Heart Rate (bpm) 60 60 60 60 120 % Time in Systole 40 % 40 % 40 % 40 % 50 % Compressor Reg Pout (bar) 1.9 7.5 0.25 7 6.5-8 Regulator Pin (mmHg) 117 88 108 62 197

Resistance Pinch Valve Signal (Volt) 1.47 4 1.38 2 1.71 Vascular Resistance (PRU) 1.08 0.1 0.95 0.22 0.63 Vascular Resist. (dyne.s.cm-5) 1439 133 1266 293 839

Mean Circulatory Pressure Added Water Volume (mL) 400 400 1100 1100 1100 Pmc (mmHg) 7 7 20 20 20

Compliance Distance of Test Plug from arterial chamber top 40mm 120mm 150mm 120mm 230mm

Arterial Compliance 1.85 2.94 1.2 2.95 0.4 Distance of Test Plug from venous chamber top 1000mm 800mm 500mm 800mm 500mm

Venous Compliance 23 7.1 11.5 7.1 4.2

Rig Configuration: The mock circulation rig was configured to operate in the

complete circulation mode to evaluate hemodynamic performance for resting and

heart failure conditions. This was achieved by opening ball valves 1 and 2, while

closing valves 3 and 4. Due to the rig limitations discussed, the exercise condition

could only be reproduced in the systemic loop. This required the closure of valves 1,

2 and 4, and opening valve 3.

Setting Compliance: Compliance values were established by positioning the test

plugs at a vertical distance from the top of each chamber. Small piping holes

through the test plugs were capped by the pressure transducers and a bleed valve.

This valve was opened to equalise chamber pressure with atmospheric.

Fluid Filling: The rig was then filled with approximately five litres of room

temperature (20oC) water. This volume filled the piping and raised the water level to

40mm above the bottom of each compliance chamber. At this point, the bleed valves

were closed, trapping a predetermined volume of air above each water level.


3-53

Setting Pmc : An additional volume of water was then added to the left and right atrial

chambers. Due to the compliance of all chambers, this volume slightly compressed

the air in each chamber. Water was added until the desired the mean circulatory

pressure was reached.

Heart Functionality: Left and right air compressors were charged to 7 Bar. The

output regulator of the left compressor was changed to reflect the degree of left heart

functionality for simulation. Left and right input regulators were also tuned to reflect

the heart condition. At this point, heartbeat was commenced. The CONTROLDESK

interface enabled the selection of heart rate and systolic period, which was

transferred to the solenoid valves.

Vascular Resistance: Vascular resistance was then tuned to reproduce the desired

hemodynamic response. Values of resistance are presented as voltage signals

between 1-5V, one representing closed and five, fully open.


3-54

0 1 2 3 41618202224

Time (sec)

Pre

ssur

e (m

mH

g)


LAP

0 1 2 3 40

50

100

Time (sec)

Pre

ssur

e (m

mH

g)


LVP

0 1 2 3 4

50

60

70

Time (sec)

Pre

ssur

e (m

mH

g)

(d) Aortic Pressure

AoPMAP

2.4 2.6 2.8 3 3.2 3.40

20

40

60

80

100

120

Time (sec)

Pre

ssur

e (m

mH

g)


LAPLVPAoPMAP


0 1 2 3 42

4

6

Time (sec)

Pre

ssur

e (m

mH

g)


RAP

0 1 2 3 40

20

40

Time (sec)

Pre

ssur

e (m

mH

g)


RVP

0 1 2 3 410

20

30

Time (sec)

Pre

ssur

e (m

mH

g)


PAPMPAP

0.8 1 1.2 1.4 1.6 1.80

5

10

15

20

25

30

35

40

Time (sec)

Pre

ssur

e (m

mH

g)

(a) Pulmonic Pressure Distribution

RAPRVPPAPMPAP

1 2 3 40

20

Time (sec)

Pre

ssur

e (m

mH

g)


RAP

0 1 2 3 40

20

40

Time (sec)

Pre

ssur

e (m

mH

g)


RVP

0 1 2 3 420

30

40

Time (sec)

Pre

ssur

e (m

mH

g)


PAPMPAP

2.4 2.6 2.8 3 3.2 3.40

5

10

15

20

25

30

35

40

Time (sec)

Pre

ssur

e (m

mH

g)


RAPRVPPAPMPAP


Figure 3-33 – Pressure distribution throughout the vascular tree for systemic (a)

and pulmonic (b) circulation` systems for resting conditions

Figure 3-34 – Pressure distribution throughout the vascular tree for systemic (a)

and pulmonic (b) circulation systems for heart failure conditions.

0 0.5 1 1.5 2 2.5 3 3.5 40

1

2

3

4

5

6

Time (sec)

Per

fusi

on R

ate

(L/m

in)

MSQSQMPQPQ

0 0.5 1 1.5 2 2.5 3 3.5 40

1

2

3

4

5

6

Time (sec)

Per

fusi

on R

ate

(L/m

in)

MSQSQMPQPQ

Figure 3-35 – Perfusion Rate for Rest Figure 3-36 – Perfusion Rate for Left Heart Failure.

0 1 2 3 4

8

10

Time (sec)

Pre

ssur

e (m

mH

g)


LAP

0 1 2 3 40

50

100

Time (sec)

Pre

ssur

e (m

mH

g)


LVP

0 1 2 3 4

80

100

120

Time (sec)

Pre

ssur

e (m

mH

g)

(d) Aortic Pressure

AoPMAP

0.8 1 1.2 1.4 1.6 1.80

20

40

60

80

100

120

Time (sec)

Pre

ssur

e (m

mH


LAPLVPAoPMAP


3-55

3.5.4 Results

Results were obtained to verify the ability of the mock loop to recreate the

hemodynamic characteristics of the natural circulatory system. Complete mock loop

results for rest and left heart failure conditions are presented for comparison, before

those obtained for exercise from the systemic loop configuration are displayed.

For each result of pressure distribution, all component pressures over a single cardiac

cycle are compared in the left graph. The three graphs on the right side detail the

individual component pressure over four seconds. Systemic and Pulmonary perfusion

rates are also displayed over four seconds.

Rest and Left Heart Failure Simulation

Pressures recorded in the systemic and pulmonic loops for resting conditions are

displayed in (Figure 3-33). A mean aortic (MAP) pressure of 96mmHg and

pulmonary arterial (PAP) pressure of 17mmHg was observed. Furthermore, LAP

varied from 8-10mmHg, while RAP oscillated about 4mmHg.

The left heart failure condition encountered an expected decline in MAP to

60mmHg, and rise in PAP to 30mmHg (Figure 3-34). LAP was also predictably

higher (20mmHg), indicating a degree of pulmonary congestion. Considerable

pressure spiking and ringing was observed in the artificial atrial and ventricular

chambers. These occurrences were attributed to water hammer surges caused by the

rigidly mounted and rapidly closing brass swing check valves. These relatively high

frequency fluctuations were digitally filtered in real-time. Furthermore, a phase

difference was observed between ventricular and arterial pressure measurements.

This was attributed to the inherent inertial effects of the mock vasculature, due to the

piping cross-sectional areas and lengths, coupled with the vascular chamber

dimensions. Additionally, the physical distance from ventricle to arterial chamber

contributed to the phase delay due to the pressure wave speed.

Systemic and pulmonic perfusion was monitored during simulated resting (Figure

3-35) and left heart failure (Figure 3-36) conditions. Perfusion was reduced from a

normal value (5.15 L/min) to those expected during LHF (2.7 L/min).


3-56

Exercise Simulation

The simulated exercise condition in the systemic loop configuration produced the

pressure distribution displayed in (Figure 3-37). An elevated MAP of 122mmHg

was observed, with systolic pressure (Psys) exceeding 180mmHg while diastolic

pressure (Pdias) was maintained at 80mmHg. Left atrial pressure was 20-18mmHg

during the ventricular filling phase, indicating a larger mean circulatory pressure. An

increase in heart rate is also observed by the reduced period of cardiac cycle.

Figure 3-38 displays the perfusion rate (9.5 L/min) observed over eight cardiac

cycles, during simulated exercise conditions.

These results presented for rest, LHF and exercise show close correlation to the

native hemodynamics. Furthermore, systemic and pulmonary flows were closely

matched in each condition, indicating the system adheres to the Frank-Starling law of

the heart.

Frank-Starling Response

The rig's adherence to Frank-Starling Law is by design, and although not

quantitatively measured was qualitatively observable during operation. Changes in

preload were observed through the clear PVC piping as changes in ventricular

chamber fluid level prior to systole. A consequent change in stroke volume was also

observed through these chamber walls. Quantitative measurement of this

phenomenon is identified as future work.

0 1 2 3 40

50

Time (sec)

P (m

mH

g)


LAP

0 1 2 3 40

100200

Time (sec)

P (m

mH

g)


LVP

0 1 2 3 480100120140160180

Time (sec)

P (m

mH

g)

(d) Aortic Pressure

AoPMAP

0 1 2 3 4323436

Time (sec)

P (m

mH

g)

(e) Systemic Venous Pressure

SVP

1.8 2 2.2 2.40

20

40

60

80

100

120

140

160

180

200

Time (sec)

Pre

ssur

e (m

mH

g)


LAPLVPAoPMAPSVP

0 0.5 1 1.5 2 2.5 3 3.5 4

0

1

2

3

4

5

6

7

8

9

10

Time (sec)

Per

fusi

on R

ate

(L/m

in)

Systemic Perfusion Rate

MSQSQ

Figure 3-37 – Systemic Pressure Distribution

(Exercise) Figure 3-38 - Perfusion Rate for systemic

circulation (Exercise)


3-57

3.6 Discussion

For effective simulation of the cardiovascular system, native heart functionality,

vasculature, pressure, perfusion and autoregulation must be replicated for all

physiological conditions. The following discussion focuses on the techniques used

and results obtained from the simulation model and physical mock circulation loop.

3.6.1 Heart Functionality

Heart functionality relates to heart rate, systolic and diastolic periods, maximum

ventricular pressure, contractility and relaxation, adherence to the Frank-Starling

Law and backflow prevention.

Heart rate, Systolic/Diastolic Periods, Max Ventricular Pressure

Comparison of ventricular performance, from the native to simulated circulation and

mock circulation, revealed similar ventricular pressure pulse wave development for

each heart function condition. Maximum ventricular pressures, pulse widths (40%

systole) and heart rates (60-120bpm) dictated by native values were replicated in the

simulated model (Appendix D) and mock loop (Table 3-14).

The pneumatic technique of introducing compressed air into the ventricle chambers

to recreate the heart’s function was successful in producing pulsatile flow in the

mock circulation loop. The 3/2 solenoid valves enabled the successful variation of

heart rate (ON/OFF rate) and systolic period (ON period), while the regulators at

ventricular entrance limited ventricular maximum pressure (pulse amplitude).

Contractility and Relaxation

The native ventricle encounters a decreasing compliance during the systolic phase.

The pneumatic technique of heart functionality reproduced this characteristic by

introducing air into an initial, pre-systolic ventricular chamber air volume above the

fluid level. According to Boyle’s law, the additional air pressure increase within the

ventricle was in proportion to the rate of air inflow. The ability to easily vary the

degree of heart function (contractility) to simulate all physiological conditions, by

simply varying input air mass flow rate was provided. This was achieved by

regulating the exit pressure from the compressors. The regulators at the ventricular

entrance were unaffected and continued to limit the maximum pressure in each

chamber. The use of a large effective area solenoid valve (26mm2) allowed for


3-58

sufficient mass flow rate into and out of the ventricle. This latter situation is

responsible for ventricular relaxation, and could be altered by placing a restriction

valve on the solenoid exhaust port. The compliance of the ventricle was sufficiently

low, and in combination with inflow air resistance, provided a sufficient rise time

(contractility) response to achieve and maintain maximum ventricular pressure.

The pressure development rise time differed in simulation model. This variation in

contractility is attributed to the difference in system order of the simulated model and

physical mock loop systems. The simulation model was confined to a first order

system, characterised by no overshoot in response to a step ventricular pressure

input. The mock loop however, featured a considerable amount of overshoot and

relatively long settling time. Inertiance of the compressed air and circulating fluid

elevates the mock loop to second order. For example, the water hammer effect of the

rapidly closing, and rigidly mounted check valves caused significant pressure spikes

within the ventricle. These spikes are dampened in the native cardiovascular system

due to the compliant nature of the valves, ventricular walls and vessels. The high

frequency pressure spikes were reduced in the mock loop by the use of a small

compliance chamber and digital low pass filter.

Frank-Starling Law

The natural heart has the ability to alter contractility in response to changes in

preload. This preload is determined by venous return passively filling the ventricle

during the diastolic phase. Passive filling ventricles are an important feature, since

the natural heart is predominantly non-sucking, and relies on atrial and venous

pressures to refill the ventricle during diastole. An increase in preload, due to

increases in alternate ventricular function or mean circulatory pressure, stretches the

ventricular myocytes thus enabling a more forceful systolic contraction. This is the

guiding principle behind the Frank-Starling Law, and was qualitatively observed in

the mock loop. An increase in ventricular preload resulted in an observably higher

fluid level within the clear walled ventricular chamber. This effectively reduced the

volume of air contained within the chamber above the fluid prior to the subsequent

systolic period. An increase in contractility resulted, since the time for pressure rise

in the smaller air volume for the same input mass flow rate is lower, and therefore

stroke volume was observed to increase. This effect is reversed for cases of reduced

preload, for which the current system has an operational limitation.


3-59

There is currently no physical barrier to prevent continued ventricular ejection in

cases of very low venous return / preload. Although the system can be tuned for this

not to occur, if preceding vascular resistance increases and thus venous return

reduces to one ventricle, continued ejection will be attempted, forcing air into the

system. A suggested future method to counteract this limitation employs a light

weight ball floating in the ventricular chamber. The ball diameter would be

sufficiently large to plug the base of the ventricle pipe during cases of low preload,

thus preventing ejection. An adequate clearance with the ventricular chamber pipe is

also required to allow vertical movement.

Backflow Prevention

The brass check valves performed the function of the various heart valves by

providing low resistance to forward flow as well as successful restriction of back

flow. However, pressure fluctuations were encountered in the mock loop ventricles

and atria immediately following respective check valve closure. These fluctuations

are not found in the simulation model, since it is a first order system and no backflow

is possible.

The fluctuations were attributed to water hammer effects caused by the sudden and

forceful closure of the valve. This effect was amplified by the weight of the swing,

its metallic nature, and hence inherent non-compliance. Additionally, a small amount

of leakage would contribute to the atrial pressure fluctuations, as the swing bounces

upon closure. In the natural heart, this pressure transferral is apparent (although not

caused by fluid transferral) as the valves are compliant and bulge into to the

neighbouring chamber when closed.

An attempt was made to dampen the ventricular pressure oscillations via two

techniques. Firstly, a digital low pass filter was placed on the ventricular pressure

recordings and secondly, a small compliance chamber (syringe filled with air) was

inserted near to the closing valves. Refinement of this latter technique by altering

syringe size (compliance) and input resistance has the potential to further eliminate

the effect. In order to reduce backflow and thus atrial fluctuation, each valve could

be replaced by plastic check valves, or ultimately mechanical type artificial heart

valves. Mounting these valves on a compliance rubber base would help alleviate the

encountered problems.


3-60

3.6.2 Pressure Distribution

Blood pressure distribution within the heart chambers and vascular network was

recreated for all physiological conditions in the simulated model and mock

circulation loop. Pressures recorded in the mock loop are compared to the native,

literature and simulated model cardiovascular pressure distribution in Table 3-15.

Table 3-15 – Comparison of Mock Circulation Loop Pressure Results

mmHg

LA

P

LV

P

LV

P E

D

AoP

MA

P

RA

P

RV

P

RV

P E

D

PAP

MPA

P

Pmc

REST

Natural 8 - 10

0 -120 8 120

/80 93 3 0 - 25 3 25

/10 15 7

(Pantalos, 2004) 10.5 -9 - 139 2 - 5 125

/68 95 - - - - - -

(Patel, 2003) - - 9 140 /90 102 - - - - - -

Mock Loop 8-10 0 - 120 5 119

/78 96 4 0 - 32 4 26

/11 17 7

Simulation 7-10 0 -120 7 113

/79 97 3 0 - 36 3 27

/11 17 7

LHF

(Pantalos, 2004) 16.3 -1 -93 14 84

/43 60 - - - - - -

(Patel, 2003) - -3 -65 18 60

/40 53 - - - - - -

Mock Loop 20 15 -80 20 70

/50 60 16 17 - 40 17 35

/25 30 20

Simulation 22 22-80 23 72

/49 59.5 16 14 - 40 16 36

/25 29.5 20

EXERCISE

Mock Loop 17 - 55

0-185 10 185

/78 121 - - - - - 22

Ventricular Pressure

Ignoring the effects of water hammer, the use of inflow regulators successfully

restricted left ventricular pressure in the simulation model and physical system to a

maximum ventricular pressure for each physiological condition.

Arterial Pressure

These pneumatically generated ventricular pressures were successfully transferred to

the arterial chambers in the physical mock circulation system. This transferral

experienced a phase delay not encountered in the simulation model or native

cardiovascular function. This delay may be attributed to a combination of the

inherent mock vascular inertiance (caused by changing pipe directions and cross-


3-61

sectional areas) which, together with arterial chamber compliance and input

resistance (check valve), influenced the time response to a step input. Further

contributions to this effect could be credited to pressure wave reflections and speed,

and the physical location of the arterial chamber.

Despite the lag, pulse pressures and mean arterial pressures were maintained, due to

the use of correct chamber compliances and exit vascular resistances. In the physical

mock loop, pulse pressure was increased from 20mmHg (LHF) to 41mHg (rest) and

107mHg (exercise). Corresponding increases in MAP from 60mmHg (LHF) to

96mmHg (Rest) and finally 121mHg (Exercise) matched those expected in each

simulated model condition.

The arterial pressure wave in the simulation model (Figure 3-21) develops as a

characteristic response to a step ventricular pressure input, while it reduces in a

fashion dictated by a single value of vascular resistance. The mock loop however

(Figure 3-33), has a distinctive rising pressure development slope, culminating in a

peak pressure before diminishing with a reducing pressure slope. Native arterial

pressure waves are also characterised by varying slopes of pressure increase and

decrease, attributed to the elastic nature and thus slight area (inherent resistance)

changing response to the pressure pulse.

Slight differences in pressure pulse developments for all vascular components,

between simulation model and mock loop results, can be attributed to the difference

in system order. The simulation was modelled as a first order system, incorporating

values of compliance and resistance in an equivalent RC circuit. However, the actual

circulation rig encounters a level of fluid inertiance (inductance) and therefore

replicates a second order RLC circuit. Additionally, but less significantly, the single

and unchanging value of inherent vascular resistance employed in the simulation

model differs from the resistance characteristic of the mock loop. The vascular flow

area of the loop does not change during each pressure pulse. However, the value of

pressure differential changes during this pulse, therefore the value of maximum flow

velocity and consequently inherent resistance value changes to a small degree.

Atrial Pressure

During diastole, ventricular preload is observed as ventricular pressure increases to

the value of atrial pressure. The Left atrial pressure in the mock loop was maintained


3-62

between 8 and 10 mmHg, for normal heart resting conditions. The atrium of the

failed left heart experienced elevated pressures (20mmHg), predominantly due to the

reduction in left heart function and increase in Pmc.

Left atrial pressure is difficult to maintain in a systemic only circulation loop that

relies on passive filling ventricles. The native RAP is lower than LAP, and since

these components are merged in the systemic only loop, either the RAP is too high or

LAP is low. The former situation requires a lower SVR than normal to provide

sufficient perfusion, since the arterial/venous pressure gradient is lower. The latter

case reduces the operation of the heart, since the atrial/ventricular pressure gradient

is lower, and ventricular filling is impaired. This may be overcome by considerably

reducing the atrial/ventricular valve resistance, or incorporating a sucking type

ventricle to create a negative pressure in the ventricular chamber to induce flow. The

mean circulatory pressure also compounds the problem, since the commencement of

the heart beat shifts fluid from the venous to arterial chambers. Therefore, the venous

pressure falls in proportion to the rise in arterial pressure and chamber compliance

values. Therefore, larger Pmc than native values is another requirement for systemic

only mock loop operation. This limitation is not apparent in the complete simulation

model and mock loop results, as the right heart and thus pulmonary venous system

feeds the left atrium. Therefore, venous pressure exceeds left atrial pressure in all

cases, thus providing sufficient venous return to the heart for passive diastolic filling.

Venous Pressure

Venous pressures were maintained at reported physiological levels due to a larger

compliance chamber. This allowed sufficient venous return to the left and right

ventricles, and helped to maintain correct atrial pressures, and consequently

ventricular preload.

Mean Circulatory Pressure

The introduction of extra fluid was successful in creating the mean circulatory

pressure. Additional fluid could then increase Pmc for cases of heart failure and

exercise.


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3.6.3 Vasculature

Faithful reproduction of the vascular resistance and compliance parameters is

essential for cardiovascular reproduction. Table 3-16 and Table 3-17 detail vascular

resistance and compliance respectively in a comparison of systemic (left) and

pulmonary (right) values between native, literature reviewed, physical circulation

loop and simulation model results for rest, left heart failure and exercise.

Table 3-16 – Systemic (SVR) and Pulmonary (PVR) Resistance Comparison

Resistance (dyne.s.cm-5) Rest LHF Exercise Left Right Left Right Left Native 1463 106 1800 160 718 (Patel, Allaire et al. 2003) 1130 - 1862 - - (Pantalos, Koenig et al. 2004) 1218 - 2023 - - Mock Loop (dyne.s.cm-5) 1439 133 1266 293 839 Simulation 1552 176 1346 216 -

Table 3-17 – Systemic (SAC) and Pulmonary (PAC) Compliance Comparison

Compliance (ml/mmHg) Rest LHF Exercise Left Right Left Right Left

Arterial Compliance Native 1-2 2-4 - - - (Pantalos, Koenig et al. 2004) 1.3 - 1.3 - - Mock Loop 1.84 2.95 1.2 2.95 0.4 Simulation 1.5 3 1.3 3 -

Venous Compliance Native 10-200 4-6 - - - (Pantalos, Koenig et al. 2004) - - 1.3 - - Mock Loop 23 7.1 11.5 7.1 11.5 Simulation 22.5 10 22.5 10 -

Resistance

Resistances were determined from values of pressure and flow between each section.

The inherent resistance of each section was fixed by the pipe dimension. However,

gate and ball valves incorporated in the arterial and venous lines could be used to

fine tune individual resistance. Furthermore, the introduction of a proportionally

controlled pinch valve between arterial and venous sections represented a variable

lumped resistance in the body (systemic) and lung (pulmonic) components.

Native and mock loop resistance values for the resting and exercise conditions show

reasonable correlation, with simulation values showing a trend of slightly higher

resistance. The left heart failure condition exhibited a lower resistance in the mock

loop and simulated model to those presented in the literature. This is also contrary to

the native cardiovascular response to LHF, which acts to increase SVR from rest.


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The discrepancy is attributed to the simulated treatment. Higher values of resistance

represent the autoregulatory response to heart failure. However, drug treatment of

heart failure relies to some degree on the reduction of SVR to values approximating

rest. This action reduces the workload on the heart, and results in increased perfusion

at the expense of mean aortic pressure (MAP). Similar results are presented in

Chapter 5 for untreated heart failure prior to LVAD insertion.

Heart valve resistance was set by the physical size of the swing check valves. Despite

the physical diameter of approximately 23mm for native heart valves, larger diameter

swing check valves were incorporated. The atria-ventricular and ventricular-arterial

valve diameters were 40mm and 32mm respectively. The larger diameters not only

account for partial blocking of the flow section by the open swing, but also its heavy

weight. This was essential to maintain passive ventricular filling, and the larger size

also accounts for the lack of atrial contraction.

Compliance

Compliance values of each chamber in the physical circulation system were

evaluated by adding volume to the section and recording the change in pressure. In

the initial condition, the volume of air trapped within each chamber dictated

compliance. However, after the pulse was initiated, the shift of fluid to the arterial

side caused this volume to decrease as pressure rose to the diastolic value. Therefore,

the actual operational value of compliance was slightly less than that measured in the

static initial condition (Table 3-14).

The technique of vertically moving the chamber test plugs was effective in easily

changing arterial and venous compliance according to the physiological condition.

Atrial and ventricular compliances were fixed by the pipe size of each chamber. That

is, a change in fluid level (volume) contained in each chamber gave rise to a pressure

head change due to the vertical displacement of this fluid level.

Compliance values of individual simulation model and mock loop components were

recreated reasonably close to the ranges found in the native circulatory system and

presented in the literature. Arterial compliances were slightly higher, to account for

the inertial effects of the mock vasculature, thus maintaining expected pulse

pressures. Systemic compliance was reduced in the LHF case, as dictated by the

native regulatory response to heart failure.


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3.6.4 Perfusion

Vascular perfusion rates are compared between Native, Literature, Simulation and

Mock Loop values in Table 3-18. Perfusion rates not only relate to total mean flow

(MSQ / MPQ), but peak arterial flow (AoQpk / PAQpk).

Table 3-18 – Comparison of Mock Circulation Loop Perfusion Results

AoQpk (mL/sec)

SQ (L/min)

PAQpk (mL/sec)

PQ (L/min)

REST Native 0 - 600 5 0 – 500 5 (Pantalos, 2004) 0 - 333 5 - - (Patel, 2003) - 5.2 - - Mock Loop - 5.1 - 5.1 Simulation 0 – 400 5 0 – 310 5

LHF (Pantalos, 2004) 0 - 183 3 - - (Patel, 2003) - 3.3 - - Mock Loop - 2.7 - 2.7 Simulation 0 - 200 2.67 0 - 190 2.67

EXERCISE (Patel, 2003) - 7.95 - - Mock Loop - 9.5 - -

Under normal and failing heart functions, perfusion rates for each physiological

condition were found to pulse with each cardiac cycle, averaging to flow rates

consistent with native and reported values for each condition. However, maximum

flow pulses recorded in the mock loop are somewhat reduced from those predicted in

the simulation model, due to the time constant of the flowmeter.

Native values of perfusion in the complete circulation rig could only be obtained for

conditions of heart failure and rest. Exercise perfusion was only observed in the

systemic only circulation rig. The lack of complete circulation rig ability to achieve

the exercise condition was a direct consequence of the inability to reduce pulmonary

vascular resistance to induce sufficient pulmonary flow for a given pressure gradient.

Values of ventricular ejection in the simulation model were in the range observed

natively, and reported in literature. This was due to the matching of aortic and

pulmonary valve resistance. The 32mm check valves, partially blocked by the swing,

therefore resulted in an equivalent area equal to that of natural heart valves.


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3.6.5 Cardiovascular Interaction and Autoregulation

The goal of the developed simulation and mock circulation rig was to closely

replicate the function of the cardiovascular system for all physiological conditions.

As previously discussed, this requires the replication of individual cardiac and

vascular function. However, their interaction and autoregulation is also of particular

importance.

In simulation and circulation rig applications, a change in cardiac function is met by

a change in vascular parameters, in order to maximise perfusion and maintain

physiological pressures. For example, in simulated left heart failure, a reduction in

left ventricular contractility necessitates an increase in vascular resistance to

maintain mAoP at the expense of flow rate. The cardiovascular system relies on the

activation of baroreceptors for this purpose, however this regulation was provided by

visual mAoP inspection and manual adjustment of SVR in the mock circulation rig

experiments. Implementing a control algorithm to replicate the baroreceptor effect is

a realistic task for future work. Furthermore, the response of the vasculature to

medical drug therapy is evident. That is, a reduction in SVR causes an increase in

perfusion at the expense of MAP.

The most notable interaction observable in the mock circulation rig is the Frank-

Starling effect. As left ventricular contractility function alters (for example due to

heart failure), left heart ejection fraction decreases. This in turn results in a lower

venous return to the right atrium, and thus less right ventricular preload. The

reduced preload negatively affects the pumping ability of the right ventricle, and

consequently right ventricular ejection is also reduced until equilibrium is reached

between right and left cardiac output. When any change in heart function is seen by

either ventricle, the corresponding ventricular functionality is passively altered by

changes in preload, which act to balance perfusion. This phenomenon was

qualitatively observed through the clear ventricular chambers in the mock loop.


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3.7 Conclusion

The techniques employed to recreate the complete human circulatory system

provided an accurate method for physiological pressure and perfusion simulation

throughout the cardiovascular network.

The developed physical mock circulation rig demonstrated the Frank-Starling

response to all physiological conditions, due to the recreation of passive filling

ventricles. Left and/or right ventricular heart function could be independently and

variably controlled, which is difficult to reproduce in an in-vivo animal setting.

Easily variable vascular parameters enable the complete mock circulation system to

recreate native hemodynamics for conditions of rest and medically treated heart

failure. However, due to the inherent resistance of the pulmonary circuit, exercise

could be simulated in the systemic loop only.

The developed SIMULINK model was based on the MATLAB demo ‘Physbe’, and

closely matched the physical circulation rig in regards to vascular pressures and

perfusion rates under all physiologically tested conditions. Variations are attributed

to the difference in the system order between model and rig. Despite the limitations,

the simulation model provided a valuable tool to sufficiently verify operation and

diagnose problems throughout the iterative mock loop design process.

The mock circulation rig can be used as a cost effective process to evaluate the

hemodynamic impact of left-, right- and bi- ventricular assist devices on the

circulatory system. Furthermore, the response of physiological controllers can be

assessed by attempting to maintain positive pressures within the ventricular

chambers. Additionally, the pulsatile environment is advantageous for device

endurance tests, although this falls outside the scope of the current design in present

form. Simple modification of rig materials would allow the use of various blood

analogues.

Although the mock loop will not replace in-vivo trials, by improving its performance,

sufficient results may be obtained to refine designs before these expensive and time

intensive trials.

Chapter 4 – Centrifugal VAD Design and Development

4-1

Chapter 4

Centrifugal VAD Design and Development

This chapter addresses the design and development of centrifugal pumps for Left,

Right and Bi-Ventricular support applications. Each pump requires the simultaneous

design of the impeller and the volute casing. Consideration for the interaction of

these components on hydraulic performance and impeller force generation must be

taken when selecting their individual design parameters. A custom VAD hydraulic

design procedure, detailed in Appendix B, was formulated as a guide to correctly

design various configurations of each individual pump component.


4-2

4.1 Introduction

Centrifugal pump design is a refined art spanning many decades. Most of the

consulted literature describes similar procedures for component design; however they

generally cater for higher specific speed applications than required of an implantable

VAD. Nevertheless, an attempt was made to integrate these methods into a design

procedure for the development of centrifugal blood pumps suitable for left, right and

bi-ventricular assistance.

Most modern turbo machinery is designed and optimised for a single operating point.

However some applications require the pump to function at a variety of operating

conditions over time. This instantaneous operating point of an implantable VAD

depends on both the level of support required for the failing ventricle, as well as the

period of the cardiac cycle. As the ventricle continues to fail, or indeed improves its

function, the output requirements of the assist device will change. Likewise, as the

natural heart alternates between systole and diastole, the changing inlet conditions

influence the VAD operating point. Therefore, the trade-off between optimal

performance at one operating point and the equally acceptable performance under a

range of operating conditions must be considered (Japikse, Marscher et al. 1997).

Performance of a centrifugal pump is conventionally described by the pump

characteristic curves (characteristic performance). However, when referring to

cardiac assist devices, the term extends to encompass the ability of the device to re-

establish cardiovascular hemodynamic parameters of pressure and flow from

pathological to normal levels (hemodynamic performance). This capability is indeed

influenced by the pump performance characteristic curves, where the gradient of the

curve determines the response of the device to changing physiological conditions

(exercise, rest, etc). A flat curve represents a large variation in flow rate for a small

change pressure head at a given rotational speed, a feature of larger exit angle

impeller vane profiles. This characteristic simplifies the physiological control of the

device by accommodating a large variation in outflow in response to changes in

vascular resistance, with little change in rotational speed.


4-3

Device operation results in fluid pressure distributions within the pump casing. When

variations in pressure exist, hydraulic thrust acts on the impeller. The characteristics

of the axial and radial components are influenced by the impeller type and volute

casing respectively. These force characteristics determine the required load capacity

of magnetic or hydrodynamic impeller suspension systems to minimise impeller

touchdown, and predict wear lifetime of contact bearings (Song, Wood et al. 2004).

The purpose of this chapter is to provide background information on centrifugal

pump design, hydraulic performance and hydraulic force generation. This

information was used to create a number of hydraulically designed components.


4-4

4.2 Background

4.2.1 Centrifugal Pump Design

The centrifugal pump is a constant flow device commonly selected for high pressure,

low volume (low specific speed) applications. Fluid enters the pump inlet axially,

whereby a rotating impeller urges the fluid radially outward. This impeller contains a

number of blades with a specific curvature, which can be fitted to a hub (back

shroud) and front shroud. The fluid is collected by the casing and guided into a

radially orientated outlet. A number of casing configurations are possible to achieve

the goal of fluid velocity deceleration and thus pressure increase.

This section describes the various components of a centrifugal pump that may be

considered for application to ventricular assist devices.

Impeller

A centrifugal pump imparts energy to the fluid via a motor driven impeller rotating

about a central axis. The application of centrifugal forces to the fluid within the

impeller is aided by protruding vanes. Variations in vane, shroud and geometrical

configuration compose the variety of impeller types.

Vane Configurations

Vane Angles and Shape The vane angle configuration determines the nature of the impeller type, i.e. forward,

straight and backward facing.

Entrance vane angles range between 15-50º (Curtas, Wood et al. 2002), while

discharge angles lie between 15-35º with a normal range from 20-25º (Stepanoff

1957; Curtas, Wood et al. 2002). Discharge angles greater than 90º result in

increased head with capacity, although at the expense of efficiency. This is suggested

to only arise by impulse action, which is not recommended when collecting the high

velocity jets and converting them into pressure (Stepanoff 1957).


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Vane Number The number of vanes on a centrifugal pump impeller influences the degree of relative

recirculation within the impeller. As a consequence, relative velocity distributions at

impeller inlet and exit are affected due to the inertia effect of frictionless particles.

Particles retain their orientation as the impeller rotates and therefore the particle fails

to turn with the impeller. Superposition of the recirculation flow to the flow through

the impeller reduces the velocity at the front face of the vane whilst increasing it at

the back face. This produces a velocity component in the tangential direction

opposite to the discharge velocity (cu2) and additional to the entrance velocity (cu1),

both of which act to reduce developed pressure head.

This relative recirculation is reduced with an increasing blade number; hence the

pump head is higher in this case. However, the blades act to partially block flow

from impeller eye to exit, thus a high blade number has the negative effect of

reducing pump capacity.

Slip While the physical inlet and outlet vane angles of the impeller influence the angle of

the input and discharge velocity respectively, they are not always coincident.

In an established flow, a body must move faster than the established velocity of flow

in order to exert any force on the liquid flowing in the same direction. I.e. The vane

must have an “impelling” action.

The relative recirculation in the vanes has an effect of decreasing the discharge angle

(β2 to β2’), while the input angle (β1) is increased (β1’). Should the blade angle be

changed (to β2’), the liquid would still lag the impeller angle and output at an angle

(of β2’). Slip is influenced by the number of vanes and the vane discharge angles.


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Shroud Configuration

There are three common impeller shroud configurations found in centrifugal pumps;

fully enclosed (shrouded), open or partially open.

Enclosed (Shrouded) Impeller An enclosed impeller design features a shroud covering the top and bottom of the

impeller vanes. This type of impeller configuration confines the fluid into the vane

space, resulting in friction losses due to fluid flow relative to the inner side of the

shroud (ω2/2g). Leakage from impeller exit to inlet is introduced in the small

clearance gap between the upper shroud and casing, while fluid stagnates beneath the

lower shroud.

Open Impeller An open impeller design improves performance of medium to high specific speed

pumps (ns=2500-6000) (Stepanoff 1957). This is a direct result of the reduction in

disc friction losses by eliminating the front shroud. Instead, this loss is replaced by a

hydraulic friction loss caused from the relative fluid flow against a stationary wall

(c2/2g).

Removing the shroud eliminates leakage between the shroud and casing; however it

introduces a new leakage from front to back side of the impeller vanes. This loss is

generally considered to be similar to that of a closed impeller, providing a minimum

working clearance is maintained. Increasing the clearance leads to a higher level of

leakage and thus a reduction in efficiency.

Partially (Semi) Open Impeller Design A partially open impeller design incorporates a number of vanes protruding from a

lower shroud (hub) only. The upper shroud is removed, resulting in leakage between

impeller vanes as well as the stagnation of fluid beneath the lower shroud.


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Inlet Volute

The fluid passage from suction nozzle to impeller eye can take on a number of forms.

The simplest are straight or curved elbow inlets. Advanced inlet designs, such as a

side suction volute, produce a level of inlet pre-swirl as a result of a baffle or splitter

(Lobanoff and Ross 1985; Japikse, Marscher et al. 1997). Most inlets are designed to

accelerate the fluid by reducing the area approaching the impeller eye, thus reducing

boundary layer losses.

Volute Casing

A volute casing surrounds the circumference of the impeller of a centrifugal pump

(Figure 4-2(a)). The impeller discharges the pumping medium directly into this

volute, which is commonly a spiral shaped flow passage, of circular or trapezoidal

cross section that gradually increases from the tongue to throat (Tuzson 2000).

The two basic functions of the volute are to collect and discharge the fluid through

the outlet, and the conversion of fluid kinetic energy (imparted by the impeller) to

pressure energy (Lobanoff and Ross 1985).

Since the fluid leaving the impeller is travelling at a considerably greater velocity

than that in the delivery pipe, this velocity reduction and thus energy conversion is

the primary objective of the volute. Further energy recuperation is performed in the

outlet diffuser (Lazarkiewicz and Troskolanski 1965). The volute does not generate

dynamic head, rather functioning to minimise losses in the energy conversion

process (Lobanoff and Ross 1985). This diffusion inefficiency accounts for the major

part of pump losses (Tuzson 2000), therefore pump efficiency depends on the

smallest loss of energy during this process (Lazarkiewicz and Troskolanski 1965).


4-8

Assumption of Flow in Volute

The condition of constant flow momentum or velocity must be assumed when

deciding the volute design technique to be employed.

The assumption of constant momentum presumes that the fluid obeys the principle of

conservation of momentum. This suggests the fluid momentum is constant from the

impeller tip to an arbitrary radius. In order to achieve this however, the velocity must

decrease since the radius increases. The result is a declining mean velocity of flow in

the volute from the volute tongue to the throat. This theory holds true for perfect

liquids, and does not take the effect of friction into account (Lazarkiewicz and

Troskolanski 1965). For a constant volute width, the liquid path in the volute takes

the form of a logarithmic spiral (Lazarkiewicz and Troskolanski 1965). .

The best modern pumps are designed for constant average velocity for all volute

sections, since it was revealed that the measured tangential velocity components of

the fluid did not follow the law of constant angular momentum (Stepanoff 1957).

This is achieved by designing the area of the volute to increase proportionally to the

increase of angle from the tongue to the throat (Logarithmic spiral).

Comparison of Constant Momentum and Velocity Methods There is little difference in pump efficiency resulting from either method; however

there are slight variations in pressure distribution around the volute, likely due to the

variation of velocity within the volute (Figure 4-1). As far as volute dimension

design is concerned, the constant velocity method is considerably simpler to

calculate (Lazarkiewicz and Troskolanski 1965).

Figure 4-1 – Volute velocity assuming constant momentum (a) and constant velocity (b)

(Lazarkiewicz and Troskolanski 1965)

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4-9

Volute Types

The entire pump capacity passes through the volute throat, while only part of the

capacity passes other volute sections. Therefore, the volute area should increase

from the tongue to the throat to accommodate discharge from the impeller periphery

(Stepanoff 1957). This is the case for the conventional single and double volute

configurations; however some instances call for a circular or concentric volute. Each

of these volute types are discussed in the following sections.

Single Volute A single volute (Figure 4-2(a)) is the most common type of casing, which increases

in cross sectional area from the tongue to throat. Single volute designs are often used

for low capacity/low specific speed applications (Lobanoff and Ross 1985).

Double Volute The double volute is either two single volutes in an opposed arrangement, or one

single volute split by a dividing plate (Figure 4-2(b)). To provide adequate discharge

flow, the total discharge area (throat area) of the resulting two volutes is comparable

to that of a single volute (Lobanoff and Ross 1985).

Circular Volutes A circular volute (Figure 4-2(c)) consists of a concentric casing surrounding the

periphery of the impeller. The circular casing should be considered when designing

pump with a specific speed between 500 – 600 (Lobanoff and Ross 1985). When

designing the geometry of a circular volute, the ratio of volute diameter to impeller

diameter should not be less than 1.15 or more than 1.2 and the width should suit the

impeller width. The capacity at the BEP is controlled by the outlet diameter, and

should be sized according to the value of throat velocity.

(a) Single Volute (b) Double Volute (c) Circular Volute

Figure 4-2 – –Volute Configurations


4-10

Specialised Volutes In some cases, the final volute design is a combination of conventional volute

designs. That is, the volute shape may initially follow the spiral volute until a certain

point where it reverts to a circular volute, or vice versa. Alternatively, the double

volutes may employ a splitter that does not reach the pump discharge nozzle, but

rather terminating at an arbitrary point within the volute. The design of such

specialised volutes is undertaken to meet the requirements of a specific application.

Volute Cross-Sectional Areas

In some cases, a correctly designed volute with smooth walls and a conical diffuser,

can obtain efficiencies in excess of 90%. This is partly due to the shape of the

volute’s cross section (Figure 4-3). These circular and trapezoidal shapes are the

most frequently used in single stage pumps (Lazarkiewicz and Troskolanski 1965).

Figure 4-3 – Types of volute cross sections


The velocity distribution within the cross section is not uniform, with a mean

velocity just 0.78 - 0.92 Vmax. This velocity ratio is lower in comparison to that in

pipe flow, mainly due to the high velocity core of the flowing fluid caused by the

rotating impeller (Stepanoff 1957).

It is advised to increase the final ideal volute design cross sectional area by 15-25%

to account for losses and boundary layers (Tuzson 2000).

When designing a volute for low to medium specific speed (< 1100), a rectangular

volute should be considered. This arrangement is not only easier to manufacture, but

the hydraulic losses are minimal over the relatively low Ns range (Lobanoff and Ross

1985). For a blood pump application the volute corners should be filleted to reduce

the potential for thrombus formation in the sharp corners.

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4-11

Diffuser

The diffuser is the transitional section of the pump from volute throat to delivery

pipe. It is an extension of the volute, often used to convert the last amount of kinetic

energy into delivery pressure (Stepanoff 1957).

If the velocity at the volute throat equals that at the exit flange, no diffusion is

necessary (Tuzson 2000). However, if the final cross section of the volute throat is

less than the bore of the discharge flange, the transition to match the size difference

should be performed by a diffuser. The angle of the transition δ (Figure 4-4) should

not exceed a designed value (7-13º), so as to avoid fluid separation from the walls

(Lazarkiewicz and Troskolanski 1965; Lobanoff and Ross 1985). Figure 4-5

displays the relation between the taper angle and the velocity of volute flow

(Lazarkiewicz and Troskolanski 1965).

Figure 4-4 – Volute design parameters (Lazarkiewicz and

Troskolanski 1965)

Figure 4-5 – Taper Angle Vs Throat Velocity


If the volute cross section is circular, the diffuser should take the shape of a truncated

cone. However, in some cases the volute cross section is not circular, rather

trapezoidal, rectangular or a more universal shape. In such cases, the mean angle of

divergence is found by relating the unique shape to an equivalent circular cross

section (Lazarkiewicz and Troskolanski 1965).

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4-12

4.2.2 Hydraulic Performance

Ventricular assist devices benefit from a flat pump characteristic curve. This feature

provides a favourable physiological response to changes in systemic resistance. That

is, smaller changes in rotational speed are required to maintain a set aortic pressure at

flow rates required for varying physiological conditions. This section discusses the

parameters that affect the hydraulic performance curve, such as impeller blade

discharge angle, and axial clearance gap.

Impeller Discharge Angle

The idealised performance characteristic curve of a centrifugal pump depends on the

value of discharge blade angle. This produces a dimensionless linear pressure

response with flow capacity when friction, separation, slip, recirculation and

blockage losses are neglected (Figure 4-6). Slip between the finite blades will

decrease the value of discharge angle, which results in a reduction of developed

head. Recirculation affects volumetric efficiency, and occurs between the impeller

and cutwater, as well as over semi-open or open type impeller blades. Friction losses

are dependant on fluid viscosity, and create boundary layers which combine with

blade thickness to create impeller blockage. These losses significantly contribute to

the actual curve, however their magnitude is difficult to theoretically predict.

Figure 4-6 - Non-Dimensional Performance for Different Discharge Blade Angles

22.5

90

112.5


4-13

Backward Curved Blades β < 90°

Backward curved blades are recommended for high efficiency, with a blade angle of

22.5° demonstrating best efficiency.

Radial Blades β = 90°

A radial blade impeller exhibits a flat performance curve up to 75 percent of BEP,

beyond which the curve is very steep (Karassik, Messina et al. 2000). This impeller

type is characterised by lower efficiency.

Forward Curved Blades β > 90°

Forward curved blade impellers have received relatively little research attention, and

as such an efficient design has not been forthcoming. This may be related to the

difficulties in realising a casing that can catch and convert the high velocity jets

leaving the vane tips into pressure, while permitting impulse action (Stepanoff 1957).

Attempts to simulate this type by reversing the rotational direction of backward

curved impellers resulted in poor efficiencies (Karassik, Messina et al. 2000).

Axial Clearance

The axial clearance between impeller top shroud and casing affects the head,

capacity and thus efficiency of the centrifugal pump. This effect is more pronounced

in semi-open type impellers with a high axial clearance to impeller width (b2) ratio.

A ratio of 0.05 is determined as 100% efficient, with a ratio of 0.6 resulting in a

reduced efficiency to 62.5% (Stepanoff 1957).

Therefore, a variation in axial clearance will affect the location (in the y-axis), but

not the gradient of the performance curve.


4-14

Volute

Single

Pump designs that integrate this volute with a constant velocity regime are generally

more efficient than the more complicated counterparts.

Double

The hydraulic performance of the double volute almost attains the efficiency of the

conventional single volute at B.E.P. However, it outperforms it considerably at off

design conditions, resulting in a more efficient design over the full range of pumping

capacities (Stepanoff 1957; Lobanoff and Ross 1985). This improvement is due to

the impeller discharging fluid into a more uniform pressure distribution (Stepanoff

1957).

The double volute arrangement is characterised as exhibiting greater efficiency at off

design points, however a slightly lower efficiency at design point to that of a single

volute is evident (Lobanoff and Ross 1985).

Circular

Circular volutes are used to improve the hydraulic performance of small, high head

or low specific speed units, but they impair high specific speed units. The efficiency

of pumps incorporating a circular volute is higher for pumps with a specific speed

less than 600, with the efficiency dropping to 95% of a conventional volute at higher

specific speeds (Lobanoff and Ross 1985).

The casing losses are not as significant in low specific speed (Ns) applications due to

a generally improved surface finish in these pumps.


4-15

4.2.3 Hydraulic Force

The impeller of a centrifugal pump encounters hydraulic forces in both radial

(perpendicular to axis of rotation) and axial (parallel to axis of rotation) directions.

These forces are composed of static (steady) and dynamic (unsteady) components

(Flack and Allaire 1984).

Radial Thrust

Radial (x,y) forces encountered by the impeller of a centrifugal pump are directed

perpendicular to the axis of shaft rotation, and are composed of a dynamic cyclic

component superimposed on a static steady state load. The static component is due

to non-uniform pressure distribution around the impeller, while whirl/diffuser

interactions form the dynamic component. Labyrinth forces contribute to both static

and dynamic elements (Guelich, Jud et al. 1987; Karassik, Messina et al. 2000).

Non-uniform flow distribution at inlet, and static rotor eccentricity, can also

contribute to steady radial forces. The latter particularly affects the contribution to

steady radial forces of an axial labyrinth. An axial labyrinth (clearance gap between

impeller and casing parallel to axis of rotation) simulates a journal bearing when the

rotor is statically eccentric. This situation effectively changes the magnitude and

direction of the resultant radial force by providing a restoring force and a tangential

force due to the impeller rotation. The impeller will therefore rotate and vibrate about

a new eccentric origin (Guelich, Jud et al. 1987). This eccentric rotor assembly may

actually whirl at a natural frequency even though the shaft may rotate well above this

speed (Adkins and Brennen 1988).

From a stability and dynamic response viewpoint, the reduction of dynamic forces

are far more important for improved industrial pump lifetime (Flack and Allaire

1984). This statement presumably emphasizes the nature of many shaft failures

resulting from fatigue fractures due to cyclic loading for extended periods. Although

static impeller forces contribute to excess shaft deflection and bearing wear

(Stepanoff 1957; Lazarkiewicz and Troskolanski 1965; Guyton 1971; Baun and

Flack 2003), the magnitudes can be factored into shaft size and bearing capacity

designs.


4-16

Static Radial

Steady radial (x,y) thrust predominantly arises from the non-uniform fluid pressure

distribution around the impeller circumference caused by asymmetric flow (Jery

1985; Guelich, Jud et al. 1987). Mismatch of absolute impeller fluid discharge

velocity and volute spiral angle also contributes to radial thrust by creating

separation at the cutwater. The magnitude, direction and trend with respect to flow

capacity depend on the operational specific speed and casing/volute type employed

(Guelich, Jud et al. 1987). Figure 4-7 describes the static radial thrust characteristics

of the common volute types investigated over all flow capacities. These trends are

examined in more detail in the following sections.

Figure 4-7 – Variations in Static Radial Thrust of Common Volute Types

Effect of Volute Type The type of volute within which the impeller operates significantly affects the

magnitudes, directions and trend characteristics of static radial force. Single volutes

encounter lowest radial force at the best efficiency point, while this condition occurs

at shut-off for circular casings. The radial force magnitudes of double volute designs

are relatively low and do not significantly vary over all operating capacities.

BEPSingle Volute

Circular Volute

Double Volute

Capacity (Q)

Radial Force


4-17

Single Volute

Single volutes provide uniform impeller discharge, and thus pressure distribution

around the impeller circumference, when operating at design conditions. This occurs

at one operating point only (Uchida, Imaichi et al. 1971; Lobanoff and Ross 1985;

Lorett and Gopalakrishnan 1986; Baun and Flack 1999). This condition theoretically

produces a balanced radial force; however a small but minimal steady state force

exists at or near the design point (Lazarkiewicz and Troskolanski 1965; Flack and

Allaire 1984; Guelich, Jud et al. 1987; Baun and Flack 1999).

A volute designed assuming constant fluid velocity at the operating point is

favourable when considering radial force, since pressure distribution is theoretically

the same in all volute sections (Stepanoff 1957). However, the static radial thrust

force characteristics are similar for volutes assuming constant momentum (Guelich,

Jud et al. 1987).

Pressure uniformity is destroyed when asymmetric flow results from operating the

pump at off-design conditions (Figure 4-8), or impeller displacement from design

centre by shaft misalignment/deflection (Stepanoff 1957; Lorett and Gopalakrishnan

1986; Guelich, Jud et al. 1987; Adkins and Brennen 1988) These situations produce

a degree of static radial impeller thrust. The direction and magnitude of this force

depends on the extent of pressure asymmetry, which is determined by the operating

point in relation to the design flow rate, and is influenced by specific speed

(Lazarkiewicz and Troskolanski 1965).

Figure 4-8 – Single Volute Fluid Velocity at Various Flow Conditions (Q).

(a) Shut-off , (b) Design Flow and (c) Above Design Flow


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4-18

Operation at Part Design Capacity

The destruction of pressure uniformity at part load in a single volute pump is

attributed to flow separation at the suction side of the tongue, resulting from a

smaller flow angle leaving the impeller to the geometric angle of the tongue. When

entering the stalled or recirculation zone (where flow is zero), little pressure recovery

occurs in first half of the volute. Therefore, deceleration occurs in the second half of

the volute, resulting in an increased pressure in this region, causing a thrust force

directed to the stalled region (Guelich, Jud et al. 1987). Radial thrust at part design

capacities generally acts toward the cutwater, shown in Figure 4-9. The actual

direction within the range presented, is dependant on the relative magnitudes of the

suction effect of separation at the cutwater, coupled with the positive pressure effect

of decelerating fluid at volute exit. The consequences of these relative vectors in

relation to specific speed are described in further detail later.

Figure 4-9 - Radial Thrust Direction at Operation Below Design Capacity

(Guelich, Jud et al. 1987)

Operation at Shut-Off

The maximum radial force theoretically occurs at zero capacity (Lazarkiewicz and

Troskolanski 1965), since impeller-volute angle mismatch and fluid deceleration is

maximum at this point. However, the maximum force practically depends on the

pump characteristic curve, since it dictates the flow rate that output pressure is a

maximum. This is evident in a study conducted by (Flack and Allaire 1984) where

maximum radial force was measured at 19% of design capacity.

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4-19

Operation Above Design Capacity

Uniformity is again destroyed at flow rates above design conditions. The separation

region shifts to the outlet side of the tongue, as a result of larger flow discharge

angles off the impeller. Therefore, pressure builds up on the inside of the tongue. In

addition to this, fluid velocity at the throat is higher than that at the impeller exit,

causing the fluid to accelerate toward the throat, and thus drop static pressure.

Radial force under these circumstances acts away from the volute tongue (Figure

4-10), and again depends on the relative magnitudes of cutwater and acceleration

vector components of force.

Figure 4-10 - Radial Thrust Direction at Operation Above Design Capacity


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Double / Split Volute

A double/split volute attempts to balance the distribution of pressure around the

impeller circumference, effectively reducing radial thrust at off design conditions of

a single stage pump (Figure 4-11) (Stepanoff 1957; Lazarkiewicz and Troskolanski

1965; Lobanoff and Ross 1985; Japikse, Marscher et al. 1997; Karassik, Messina et

al. 2000).

Force minimisation is attributed to a second stalled region (Figure 4-11) created at

180o from the original cutwater, which considerably increases flow symmetry and

pressure distribution at off design conditions (Guelich, Jud et al. 1987). Volutes of

this design are often termed double or split type, though each description features a

subtle difference. That is, the double volute has a constant outer channel area, while

the splitter has an expanding outer channel area.

Non-uniformity in the double volute casing develops from different fluid resistances

of inner and outer volute channels. The double volute arrangement is therefore

sensitive to the geometrical tolerance of both cutwater and volute passages, and can

easily destroy pressure distribution symmetry (Stepanoff 1957; Guelich, Jud et al.

1987). Consequently, the double volute will greatly reduce, but not completely

eliminate radial thrust (Lobanoff and Ross 1985).

Figure 4-11 – Flow Separation Regions in a Double Volute at Part Capacities


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4-21

Modified Concentric / Volute

A number of techniques have been investigated to reduce static and dynamic radial

thrust by modifying the volute casing. Cutting back the cutwater or tongue, and

forming a concentric volute over the first 90 degrees, eliminates the stalled region at

the volute tongue at part loads since the volute tongue angle is reduced and better

matches the discharge angle from the impeller (Figure 4-12).

This technique results in a reduced thrust force at part loads (Guelich, Jud et al.

1987) at the expense of efficiency, since recirculation is increased (Uchida, Imaichi

et al. 1971). Extending the concentric volute to 270 degrees before reverting to a

volute type casing to discharge also produces a marked reduction in radial force

relative to a single volute (Lazarkiewicz and Troskolanski 1965).

Figure 4-12 – Modified Concentric Volute


Concentric

Symmetrical pressure distribution and thus minimal radial force occurs at zero flow

(shut-off) conditions in a concentric volute. This uniformity is destroyed as the

capacity increases, leading to increased radial loads and casing losses. Non-uniform

pressure is therefore often realised at BEP (Lobanoff and Ross 1985; Karassik,

Messina et al. 2000), which leads to a maximum radial force (Stepanoff 1957). A

reduction of radial load may be realised by adding material to the cutwater, which

effectively reduces fluid recirculation at the tongue (Lobanoff and Ross 1985).

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4-22

Static Radial Force Calculation The radial force encountered by the impeller of a centrifugal pump is often difficult

to theoretically predict. Therefore, many authors have derived specific formulas

based on experimental data to aid in the initial estimation of magnitude (Stepanoff

1957; Lazarkiewicz and Troskolanski 1965; Lobanoff and Ross 1985; Lorett and

Gopalakrishnan 1986; Karassik, Messina et al. 2000). These calculations are not

common, and no standard method has been accepted (Japikse, Marscher et al. 1997).

Radial thrust direction is even more difficult to predict exactly, however general

trends are identified in relation to output capacity.

Magnitude

The magnitude of radial thrust varies for each volute type. Equations for radial thrust

calculation (detailed in Appendix C) are provided by:

1. Stepanoff, (Stepanoff 1957)

2. Lobanoff/ANSI/HI, (Lobanoff and Ross 1985; ANSI/HI 1994)

3. Karassik (Karassik, Messina et al. 2000)

Each technique is based on knowledge of the output pump head, which acts on the

impeller peripheral area (including shrouds). This resultant force is multiplied by an

experimentally determined thrust factor ‘K’, which varies in magnitude from shut-

off, to above design conditions, at a trend dependant on the volute type.

These equations are implemented to predict the radial force expected in the single

sided LVAD condition. Comparison to the experimental force results presented later

in this chapter will reveal the most appropriate equation for use in future centrifugal

VAD designs.

Direction

The exact direction of radial thrust throughout the full capacity range depends on the

pump specific speed (Ns). However, resultant force generally acts toward the

cutwater during part flow, and away during high flow for all specific speeds in a

single volute casing.

The magnitude and general direction of radial thrust depends on the operating

capacity relative to the design capacity. This radial thrust vector can be further

analysed into two individual vector components. The first represents the force due to


4-23

separation/increased pressure at the cutwater, while the second vector is influenced

by the degree of acceleration/deceleration upstream from the volute throat. The

resultant vector is therefore influenced by the relative magnitude of each component.

The individual component magnitudes are dependant on specific speed. The relative

impeller-volute angle mismatch has a larger range with higher Ns pumps operating at

part flow conditions. Therefore, the separation vector is dominant and is most

influential in radial thrust direction. On the other hand, the impeller-volute match

angle is smaller in pumps of low Ns, and thus a lower range of mismatch can occur

while operating at off design. This results in a dominance of the acceleration/

deceleration vector component, since the degree of separation at the cutwater is

limited.

This observation is supported by the results displayed in Figure 4-13. Thrust

direction for higher Ns pumps at shut off is dominated by cutwater separation, and

thus lies in the 2nd quadrant. The direction for low Ns pumps has a reducing tendency

to act toward the cutwater, and thus projects into the 3rd quadrant. Similarly, for

conditions above design, the higher Ns pumps are dominated by the pressure build up

at cutwater, while a lower Ns is influenced by volute exit acceleration.

Figure 4-13 - Radial Thrust Direction for Single Volutes

(Karassik, Messina et al. 2000) (Agostinelli, Nobles et al. 1960; Lazarkiewicz and Troskolanski 1965)


4-24

Dynamic Radial

Unsteady radial forces arise from cyclic fluctuations that occur at sub-synchronous,

super-synchronous or synchronous frequencies relative to the impeller speed.

The tolerance of an impeller’s manufacture can lead to unbalances due to pressure

fluctuations created from a blade of different discharge angle. This hydraulic

imbalance has the same frequency as the impeller rotational speed and is thus termed

synchronous (Flack and Allaire 1984).

Low frequency or sub-synchronous fluctuations that result from separation of flow at

part loads will increase at rotating stall (shut-off) operating conditions, and can be

destabilising (Flack and Allaire 1984; Guelich, Jud et al. 1987; Karassik, Messina et

al. 2000).

High frequency (super-synchronous) fluctuations are generated at the blade passing

frequency due to the finite blade number.

Impeller/diffuser interactions contribute to the dynamic force when the impeller

vibrates or whirls about the origin. This vibration also causes a variable pressure

distribution about the impeller, since the distance from tip to casing varies. However

the magnitude is significantly lower than labyrinth forces and is often neglected

(Guelich, Jud et al. 1987).

Dynamic forces are difficult to predict, and are significant at 30% design capacity

with a frequency in the order of 1/10th rotational speed (Flack and Allaire 1984).


4-25

Axial Thrust

The thrust in the axial (z) direction of a centrifugal pump impeller results from a

difference in pressure distribution acting on the hub and top shroud faces. Many

factors contribute to axial thrust, such as impeller diameter, labyrinths, pump head,

and momentum change through the impeller. Accurate prediction of axial thrust is

therefore only possible with supplemental experimental data (Guelich, Jud et al.

1987). Impeller type also has a considerable effect on axial thrust, with the

magnitude of force decreasing for the semi-open, open and closed type respectively

(Figure 4-14). A number of methods for reducing axial thrust are implemented

industrially, including radial ribs, balance holes and double suction type impellers.

These techniques are sometimes used in blood pump design, with the additional aim

to improve secondary flow beneath the impeller and reduce thrombosis formation.

Figure 4-14 – Variations in Static Axial Thrust in all Volute Types

Axial Force

Capacity (Q)

Closed Type

Semi-Open Type

Best Efficiency Point


4-26

Static Axial

The static axial thrust developed in single suction centrifugal pumps is due to

asymmetry between the high discharge pressures acting on the underside of the

impeller, and the suction pressure acting on the impeller front (Stepanoff 1957). The

pressure encountered by the underside of the impeller is maximum at the periphery,

and reduces as the axis of rotation is approached, due to the whirling effect of the

liquid (Lazarkiewicz and Troskolanski 1965).

Impeller Types

Closed

The closed impeller produces minimal axial thrust as back shroud pressure is

countered by front shroud pressure (Figure 4-15).

Figure 4-15 – Variations in Static Axial Thrust in an Closed Type Impeller

Semi Open

Semi-open impeller types tend to produce a higher value of axial thrust, since back

shroud pressure is only partly opposed the diminishing pressure acting on the top of

the hub (Figure 4-16) (Stepanoff 1957; Lazarkiewicz and Troskolanski 1965;

Karassik, Messina et al. 2000).

Figure 4-16 – Variations in Static Axial Thrust in a Semi-Open Type Impeller

Open

Fully open impellers produce a somewhat larger axial force than closed types, but

less force than semi-open types (Karassik, Messina et al. 2000).


4-27

Balancing Axial Thrust Techniques implemented industrially for balancing axial thrust are usually

undertaken at the expense of hydraulic efficiency (Japikse, Marscher et al. 1997).

Although these methods may achieve the desired reduction in axial thrust, their

suitability for use in blood pumps should be carefully considered.

Balancing Holes

To attempt to equalise the asymmetric axial pressures, four to eight balance holes are

drilled between the inlet blades through to the underside of the impeller (Figure

4-17(a)). This allows communication of fluid from impeller back to front, however

equilibrium is not reached in practice owing to the resistance to flow of the holes.

Additionally, this technique can potentially reduce efficiency due to loss of delivery

through the holes, coupled with a flow disturbance at the inlet (Stepanoff 1957;

Lazarkiewicz and Troskolanski 1965).

This technique has the added benefit of providing a degree of washout beneath the

impeller of a blood pump. A reduction in stagnation is observed and thus the

incidence of thrombus formation is somewhat reduced (Nishida, Yamane et al.

1998).

Figure 4-17 – Techniques for Balancing Axial Thrust

(Japikse, Marscher et al. 1997)

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4-28

Radial Ribs

Installing four to six radial ribs to the underside of the impeller (Figure 4-17 (b)) is

the most effective way of reducing the axial thrust generated during operation

(Stepanoff 1957). This technique effectively reduces the pressure within this cavity,

by causing the fluid to rotate at the impellers speed. This results in a reduced upward

axial force and therefore reduced overall axial force. However this is at the expense

of an additional power requirement and thus loss of efficiency (Stepanoff 1957;

Lazarkiewicz and Troskolanski 1965; Japikse, Marscher et al. 1997).

Radial ribs (secondary vanes) have been investigated for use in blood pumps, again

in an attempt to improve flow patterns beneath the impeller shroud to reduce

thrombosis (Ohara, Makinouchi et al. 1994; Ichikawa, Nonaka et al. 2002).

Double Suction

By introducing an additional inlet (Figure 4-17 (c)), axial forces on the impeller are

theoretically balanced due to symmetry (Stepanoff 1957; Lazarkiewicz and

Troskolanski 1965; Lobanoff and Ross 1985). However, a residual force is

apparent in most practical applications (Stepanoff 1957; Japikse, Marscher et al.

1997; Karassik, Messina et al. 2000), and in fact is somewhat encouraged (Lobanoff

and Ross 1985). A small axial unbalanced is recommended to lightly load the

impeller bearings in one direction, and therefore prevent floating of the rotor.

Unbalance may arise from non-uniform fluid entrance to the impeller eye and

different piping conditions near the impeller eye (Karassik, Messina et al. 2000).

A double suction impeller will eliminate stagnation beneath the impeller of a single

sided centrifugal blood pump. However, the requirement for an extra inlet may

increase device size, and therefore must be carefully designed to maintain a small

device.


4-29

Static Axial Force Calculation

Magnitude

The magnitude of axial thrust varies for each impeller type. Equations for axial

thrust, detailed in Appendix C, are provided by;

1. Lazarkiewicz (Lazarkiewicz and Troskolanski 1965)

2. Stepanoff (Stepanoff 1957)

3. Lobanoff (Lobanoff and Ross 1985)

4. ANSI/HI (ANSI/HI 1994)

Each method for axial thrust calculation accounts for the contribution of fluid

momentum and pressure distribution underneath and above the impeller. The

calculations assume the geometry of the impeller and characteristic performance

pressure is known. Values of axial thrust at all capacities can be inferred from this

data. Finally, the weight of the impeller is generally not included, and therefore must

be added to the value of axial thrust for pumps operated in the vertical position.

Direction

Three force vectors, that each act in the direction of the impeller rotational axis;

resolve to produce an axial force toward the suction eye. The component generated

from pressure distribution on the impeller bottom contributes to this direction,

opposed by the pressure acting on the top shroud as well as the small contribution of

the fluid momentum component, generated from the change of fluid direction.

Dynamic Axial

Dynamic forces encountered in the axial direction increase in the recirculation zone

encountered at low capacity, and cause cyclic stress to the shaft and bearing. Since

the axial momentum is quite low for radial impellers, slight changes in capacity do

not result in significant dynamic forces (Karassik, Messina et al. 2000).

Additionally, most practical applications operate under relatively static pressure and

capacity, and thus dynamic axial thrust is not as significant as dynamic radial thrust.

The pressure and capacity of an implantable VAD is not static when supporting a

failing ventricle. However, since the operating conditions change with each heart

beat, the frequency is low. Therefore, the axial force alternates between static values

determined by the operating points during the systolic and diastolic periods.


4-30

Additional Impeller Forces

Gravity / Buoyancy

The mass of the impeller is affected by gravitational acceleration, which produces a

static force on the rotor. The magnitude of the static force is directly related to the

mass of the rotor, since the gravitational constant remains unchanged. For example, a

typical rotor mass of 50grams would produce a force of 0.49N.

To counter gravitational effects, the density of the rotor can be matched with that of

the surrounding fluid. For example, a rotor would become buoyant in a blood

environment should its density equal 1050 kg/m3.

Inertia

The rotor is exposed to inertial effects as the pump changes direction, which occurs

as the patient ambulates, or indeed as the remaining cardiac function alternates

between systole and diastole. The impeller will continue its motion until an opposing

force produced by the bearing system acts on it.

The magnitude of this force is dependant on the acceleration/deceleration of the

pump in the patient, which is expected not to increase above 2-3 G in everyday

activity (Tansley, Cook et al. 2000). Again, with a typical rotor mass of 50 grams, a

restoration force of 1.0 - 1.5 N would be required in this case.

Hydrostatic Forces / Damping

Hydrostatic forces are also present in a centrifugal pump. This type of force acting

on the impeller is predominant in labyrinth (clearance) sections. Any movement of

the impeller from its central axial or radial position will be met by an opposing

hydrostatic force. The magnitude of this force is dictated by the fluid properties, and

is dominated by viscosity. Some degree of damping also results to restrict the

dynamic motion of the rotor.


4-31

Hydrodynamic Forces

Hydrodynamic force relates to the restoration forces produced in the clearance gaps

of a fluid film bearing. Impellers that incorporate tapered wedges or journal sections

benefit from the production of these forces; however the magnitude is influenced by

the relative surface velocities and fluid viscosity.

When employing a hydrodynamic bearing, the gravitational force mentioned

previously benefits the stability of the system by introducing a bias force. The

direction of the force should coincide with the direction of the fluid film.

Magnetic Negative Stiffness

Some third generation centrifugal VAD’s employ a motor that magnetically couples

a stator core and coils to magnets embedded within the impeller. This coupling

produces an attractive force from rotor magnet to iron stator core, which is amplified

as the rotor moves from the neutral position, and thus destabilises the system. The

magnitude of the destabilisation force depends on the magnetic material strength and

distance to stator (air gap), and must be overcome by the bearing system. For

example, a negative stiffness of 20N/mm produced in a magnetic bearing application

would require a restoration force of 20N should the impeller displace 1mm from its

neutral position.

Gyroscopic Forces

Gyroscopic forces are encountered by the rotor of an implantable centrifugal blood

pump when a patient bends or turns. The magnitude of these forces depends on the

impeller operating rotational speed, impeller moment of inertia, and rate of

turn/bend. For a typical impeller diameter (50mm) and mass (50 grams) rotating at

2000 rpm while precessing at 1 revolution per second, a restoring toque of 0.02 Nm

is expected, which roughly translates to a force of 0.85 N at the outer diameter.


4-32

4.3 Design Procedure and Parameter Calculation

The procedure used for each VAD application is outlined in Appendix B. This

procedure outlines the various design parameters assumed or calculated from graphs

and equations given by numerous pump design authors. The procedure was

reproduced in MATLAB to enable efficient calculation of parameters.

Impeller design is initially undertaken to determine the operational rotational speed,

diameter and specific speed. The design then focuses on the impeller vane outlet

velocity triangle calculation, which is influenced by head and capacity co-efficient

selection. This determines the blade discharge angle and absolute fluid velocity

leaving the impeller. Inlet vane angles are also selected to minimise the degree of

pre-whirl. Finally, the impeller profile is completed by designing the transition from

inlet to outlet vane angle whilst maintaining an acceptable diffusion ratio so as not to

induce separation.

The volute casing design is then undertaken using parameters generated from the

impeller design procedure. In particular, throat velocity and thus area is determined

from the absolute discharge velocity, while the volute angle of spiral is selected to

match the angle of this fluid velocity leaving the impeller. The volute design is

continued by selecting an appropriate width and cross section, before completing the

design by creating the single, double or circular volute profile.

The entire centrifugal pump design is then completed by creating the flat inlet volute.


4-33

4.3.1 Design Conditions and Fluid Properties

The following tables describe the fluid properties (Table 4-1), the design conditions

of pressure (Table 4-2), and flow rate (Table 4-3) required of each assist application.

Table 4-1 – Fluid Properties

Table 4-2 – Design Output Pressure

LVAD Bi-LVAD Bi-VAD

Left Right Left Right

mmHg 100 100 100 100 20 KPa 13.33 13.33 13.33 13.33 2.67 meters[Blood] 1.29 1.29 1.29 1.29 0.26 Feet[Blood] 4.23 4.23 4.23 4.23 0.85

Table 4-3 - Design Flow Rates

LVAD Bi-LVAD Bi-VAD

Left Right Left Right

L/min 5 2.5 2.5 5 5 m3/sec 8.33E-05 4.17E-05 4.17E-05 8.33E-05 8.33E-05 GPM 1.32 0.66 0.66 1.32 1.32 ft3/s 2.94E-03 1.47E-03 1.47E-03 2.94E-03 2.94E-03

As identified in Chapter 3, average systemic and pulmonary pressures of 93mmHg

and 15mmHg respectively are required to overcome normal systemic and pulmonary

vascular resistances, and ensure sufficient organ and tissue perfusion of 5L/min.

However, continuous flow blood perfusion results in vasoconstriction, which

increases vascular resistance. To ensure the pump has sufficient capacity to deliver

the required flow, slightly larger values of 100mmHg and 20mmHg were selected as

the design operating points.

Due to the energy contribution of the failing heart, the pump may not always be

required to deliver exactly these conditions. However, the contribution of the device

is impossible to predict for each individual patient. Furthermore, the pump will

actually operate about these specified conditions during systole and diastole,

effectively averaging the output to deliver these requirements.

Water Blood Density (ρ) 1000 kg/m3 1056 kg/m3

Viscosity (µ) 1.15x10-3 Pas 4x10-3 Pas Kinematic Viscosity (ν) 1.15x10-6 Pas 3.8x10-6 Pas


4-34

Design parameters are presented for three pump configurations. The LVAD and Bi-

LVAD configurations are for the left ventricular support condition, and thus require

an output pressure of 100mmHg. The former is a single sided impeller device, which

therefore requires all 5L/min of flow to pass through the impeller vanes. The latter

configuration is a double sided impeller device, which requires the entire 5L/min of

flow to be diverted between each set of impeller vanes. The final configuration

attempts to address bi-ventricular support with a double sided impeller. The left side

of the impeller must deliver the systemic circulation hemodynamics, and is identical

to the single sided LVAD support condition. The right side however, must deliver the

hemodynamic conditions required of the pulmonary circulation. Since both

chambers of the device are effectively operating in series, the entire 5L/min of flow

must pass through each set of impeller vanes.

4.3.2 Pump Constants from Similitude

The following parameters are most important for centrifugal pump design, and are

used to compare the relative performance of all centrifugal pumps.

Reynolds Number (Re)

The Reynolds number is a dimensionless parameter used to relate viscosity, impeller

rotational speed and diameter (Equation [ 4-1 ]). The Reynolds number is used to

ascertain a laminar or turbulent regime, with the transition occurring around 1.0e5 for

pumps. The Re for each VAD is given in Table 4-4.

νω

4Re

2D=

[ 4-1 ]

Table 4-4 – Reynolds Number (Re)

LVAD Bi-LVAD Bi-VAD

L and R Left Right

Re 3.3 e4 3.3 e4 3.3 e4 6.6 e3

Specific Speed (Ns)

The specific speed (Ns) or type number of a centrifugal pump is a dimensionless

number that describes the relationship of pump speed, capacity and head (Equation [

4-2 ]). Specific speed (Table 4-5) may be described using imperial or metric scales.


4-35

43

21

.

H

QNN s = [ 4-2 ]

Table 4-5 – Specific Speed (Ns)

LVAD Bi-LVAD Bi-VAD

L and R Left Right

ns(US Units) 750 530 750 2507

ns(Metric Units) 0.2744 0.20 0.2744 0.916

Specific Capacity (qs)

The specific capacity (qs) is defined as the volume rate of fluid (Q[ft3/s(m3/s)] ) per

unit rotational speed (n[rad/s]) with a unit outlet radius (r2[ft(m)]), and remains

constant for similar impellers (Equation [ 4-3 ]). Values are displayed in Table 4-6.

3nrQqs = [ 4-3 ]

Table 4-6 – Specific Capacity (qs)

LVAD Bi-LVAD Bi-VAD

L and R Left Right

qs 0.0264 0. 0264 0.0264 0.298

Specific Head (hs)

The specific head (hs) is defined as the input energy per unit mass

(g[ft/sec2](m/s2),H[ft](m)) per revolution (n[rad/s]) and with an impeller of unit

diameter (D[ft](m)). It again remains constant for similar impellers (Equation [ 4-4 ])

with values for each VAD given in Table 4-7.

22rngHhs = [ 4-4 ]

Table 4-7 – Specific Head (hs)

LVAD Bi-LVAD Bi-VAD

L and R Left Right

hs 0.51 0.51 0.51 0.504


4-36

4.4 Design Detail

A number of pump impeller and volute configurations were designed for each left,

right and bi-ventricular assist application, for the purpose of assessing hydraulic

performance and impeller force characteristics. The following section discusses the

details of these configurations.

Each design was guided by the customised procedure outlined in Appendix B, which

also details the design parameters and all impeller/volute geometries. Future

iterations are recommended to best match the design with the constraints demanded

from the blood pump application.

4.4.1 LVAD

A single sided LVAD was designed primarily as a control for performance and

impeller force characteristic comparison with each subsequent VAD design. This

particularly involved the examination of the effect of exit blade angle on the

performance curve, and the variation of impeller hydraulic forces within each volute

configuration.

Table 4-8 describes the variety of impeller and volute configurations used in the

LVAD experiments. Each volute tested with the backward and radial blades were

designed with respect to those impellers. However, the forward facing impeller was

tested with the volute designed for the backward facing blade impeller, since no

design parameters were available for the design of a single volute incorporating a

forward facing impeller.

Table 4-8 – LVAD Impeller and Volute Configurations

Impeller Volute 22.5 º Backward Facing Single

Double 90 º Radial Circular

Single 112.5 º Forward Facing Single

Isometric and cross-sectional views of an LVAD pump configuration modified to

allow for appropriate testing procedures are shown in Figure 4-18.


4-37

(a) Isometric View (b) Cross Sectional View

Figure 4-18 – LVAD Pump

4.4.2 RVAD

A single sided RVAD was designed to provide the ability to test the independent

operation of RVAD impeller and volute designs.

The aim of this design was to determine the pump performance within the various

volute configurations.

Table 4-9 describes the impeller configuration designed for single RVAD testing,

with corresponding volute configurations.

Table 4-9 – RVAD Impeller and Volute Configurations

Impeller Volute 22.5 º Backward Facing Single

Circular

Isometric and cross-sectional views of a RVAD pump configuration modified to



Figure 4-19 – RVAD Pump


4-38

4.4.3 Bi-LVAD

The double suction Bi-LVAD was designed for the purpose of performance

characteristic and impeller force distribution comparison to the conventional single

sided LVAD.

The aim of this design was to assess the change in performance characteristics as

well as impeller force distribution, particularly in the axial direction.

Table 4-10 describes the impeller and volute configuration designed for Bi-LVAD

testing.

Table 4-10 – Bi-LVAD Impeller and Volute Configurations

Impeller Volute

22.5 º Backward Facing Single

Isometric and cross-sectional views of a Bi-LVAD pump configuration modified to



Figure 4-20 – Bi-LVAD Pump


4-39

4.4.4 Bi-VAD

The Bi-VAD pump design is a combination of LVAD and RVAD pump designs

joined in a single device. As such, the design parameters for each side of the device

are identical to those for the LVAD and RVAD case.

The aim of this design was to determine the combined effect of LVAD and RVAD

configurations on performance characteristics in the Bi-VAD setting.

Table 4-11 describes the variety of impeller configurations designed for the Bi-VAD

application, with corresponding volute configurations.

Figure 4-21 illustrates the Isometric and cross-sectional views of a Bi-VAD pump

configuration modified to allow for appropriate testing procedures.


Figure 4-21 – Bi-VAD Pump

Table 4-11 – Bi-VAD Impeller and Volute Configurations

Impeller Volute Left Right Left Right

22.5º Backward Facing 22.5 º Backward Facing Single Single 90º Radial Circular Circular


4-40

4.5 Discussion

The customised centrifugal pump design procedure combined many author

recommendations relevant to the cardiac support device application. The generated

MATLAB code improved the design task by efficiently calculating parameters for

impeller and volute construction.

Specific speeds (Ns) of 750 and 2507 were determined for the LVAD and RVAD

applications respectively, based on impeller diameters of 50mm and 22.3mm, and a

common rotational speed of 1921 rpm.

Each impeller incorporated six vanes in a semi-open or closed shroud configuration.

Vane angles of 22.5° (backward), 90° (radial) and 112.5° (forward) were designed

for evaluation in single, double and circular volute configurations. Although

parameters for the radial impeller were selected for 90°, the actual discharge angle

was 78°. This was due to the incorporation of a correct inlet angle (to reduce pre-

swirl) and maintenance of straight radial vanes. Parameters for the 112.5° were

interpolated from relevant graphs, as this discharge angle is not accounted for in

conventional pump design.

The double volute was a splitter type, while the base circle of the circular volutes

was 1.15 times the impeller diameter. Each volute, designed for constant fluid

velocity, incorporated a rectangular cross section with filleted edges. The width of

this cross section was maintained according to recommendations, along with a

matched volute throat area to each impeller. This resulted in a mismatch of volute

spiral angle and absolute impeller discharge angle, but was assumed negligible due to

the low specific speed. Finally, an inlet volute was designed for improved anatomical

compatibility, while the outlet diffuser angles were maintained below 13°.


4-41

4.6 Conclusion

The customized centrifugal pump design procedure facilitates the efficient and

timely design of VAD iterations. Incorporating the procedure into MATLAB code

enables the ability to easily vary the design requirements to produce a set of

parameters that influence impeller/volute development.

The procedure was used to design the double suction Bi-LVAD pump, as well as

single suction LVAD and RVAD pumps. These latter designs couple to form the Bi-

VAD pump, and were constructed to independently test various impeller and volute

configurations in each assist application.

Since the Bi-LVAD device requires identical pressure from each outlet, the vanes of

the double impeller are symmetrical in nature. To produce the variation in pressure

required from each output of the Bi-VAD device, the impeller vane diameters of

each side are different in length, as dictated by the LVAD an RVAD designs.

Chapter 5 - VAD Experimental Evaluation

5-1

Chapter 5

VAD Experimental Evaluation

The experimental evaluation of the VAD hydraulic designs presented in chapter 4

specifically investigates the effect of impeller (semi-open/closed) and volute (single,

double, circular) components for each ventricular assist application. This was

conducted within a unique test facility, to assess performance and isolate impeller

hydraulic forces while inserted into the complete mock circulation loop (Chapter 3)

configured for non-pulsatile and pulsatile operation.

The results presented in this chapter are used to recommend appropriate impeller /

volute configurations for designs incorporating magnetic, hydrodynamic and

mechanical bearings, to minimise impeller touchdown and increase device lifetime.

5.1 Introduction

The performance characteristics of a new prototype VAD are generally tested by

setting pump rotational speed and varying circuit resistance in a non-pulsatile

circulation loop (Tansley, Vidakovic et al. 2000; Masuzawa, Ezoe et al. 2003).

However, VAD performance also includes the ability of the device to assist a failing

heart to maintain sufficient cardiac output and native physiological pressures

(Koenig, Pantalos et al. 2004). For accurate testing, prototype pumps should be

subsequently inserted into a pulsatile circulation system, such as the in-vitro pulsatile

mock circulation loop presented in Chapter 3, or an in-vivo animal model.

Previous investigations into rotary blood pump forces are limited in number.

Experimentally unverified CFD models were used to evaluate forces at a single, non-

pulsatile operating point (Qian and Bertram 2000), thus effectively neglecting failing

heart pulsatility. Magnetic coupling techniques have been used to determine impeller

axial forces only (Takami, Makinouchi et al. 1997) while magnetic bearings were

calibrated to determine radial force (Allaire, Kim et al. 1996). Therefore, a unique

method was required to measure the magnitude and direction of impeller hydraulic

thrust during non-pulsatile and pulsatile device operation.


5-2

Figure 5-1 – Experimental Process Diagram

Refer Figure 5-2


5-3

The purpose of this investigation was to assess the performance (Characteristic and

Hemodynamic) of the hydraulically designed impeller/volute configurations

presented in Chapter 4, while recording static hydraulic force magnitude and

direction. The effect of impeller exit blade angle and variation of axial clearance gap

on characteristic hydraulic performance was investigated. The LVAD configuration

was inserted into the mock circulation loop configured for left heart failure, and its

ability to reduce pulmonary congestion and restore perfusion levels from

pathological to normal levels is demonstrated. The influence of failing heart

pulsatility on the magnitude and direction of hydraulic forces is therefore established,

by comparing these results to those obtained during non-pulsatile operation.

The results initially aid the selection of the best impeller/volute configuration for the

magnetically suspended Bi-LVAD and Bi-VAD designs. However, the results may

be also applied to VAD designs incorporating hydrodynamic and mechanical type

bearings. The incidence of impeller touchdown can then be minimised by

accommodating sufficient bearing load capacity to counter the force magnitudes.

5.2 Experimental Method

The following overview of the experimental configuration, used to record pump

performance and force data under non-pulsatile and pulsatile conditions, makes

reference to the process diagram displayed in Figure 5-1.

The experiments were undertaken with software and hardware installed on a 700

MHz Celeron PC running Windows 98.

The operation and capture of mock loop and VAD performance results was

undertaken with CONTROLDESK software. The systemic and pulmonary vascular

resistances, heart rate and period in systole were determined in this environment for

mock loop operation, while a value for VAD rotational speed was also set. This

information was passed to the SIMULINK model. This model incorporated blocks

for all aspects of the real mock circulation loop and VAD system. Before the

commencement of the experiment, this model was uploaded to the dSPACE card

using Real-time Workshop Build (RTW) and the Target Language Compiler (TLC).

This process provided the ability to create a link between the CONTROLDESK

interface and the experimental rig through the MATLAB Real Time Interface

toolbox. Signals could now pass in real-time from CONTROLDESK to the dSPACE


5-4

card through the MATLAB based SIMULINK model and RTI MATLAB toolbox.

An interface box transmitted the SVR, PVR, HR, Percent in Systole (%sys) and

motor speed signals from the dSPACE card to the mock circulation rig and VAD

respectively. These signals commenced and dictated the experiment operation. SVR

and PVR signals determined the actuation of the proportional control valves, each

supplied with 12VDC/2A power. HR and %Sys signals respectively controlled the

rate and period of actuation of the solenoid valves, each supplied with 24VDC

power. The RPM set signal fluctuated between the dSPACE restricted 0-10V

depending on the desired set motor speed. The signal was input into a custom 2.4

gain voltage amplifier, supplied with +/- 26VDC power, to amplify the signal to the

0-24V range of the motor.

Experimental performance was evaluated using pressure transducers and flow

meters, supplied with +12VDC supply and 240V supply respectively. The

instrumentation produced a voltage signal that indicated the actual value, which was

passed to the dSPACE card via the interface box. Voltage signals from the motor

encoder and amplifier were also returned to indicate motor speed and current, with

the latter signal calibrated to return 100mV per amp. These signals returned to the

CONTROLDESK interface through RTI, MATLAB and SIMULINK model, for

real-time capture and monitor display.

Code was written and executed in MATLAB simultaneously with the performance

parameters for the capture of impeller force data. The code utilised the MATJR3.dll

driver to communicate with the JR3 force transducer via the JR3 DSP ISA card. This

card receives voltage and current signals from the force transducer and resolves them

into force or moment values for each degree of freedom. This data is then available

for software recording/manipulation in the form of a 6x1 matrix.

5.2.1 Prototype Pump Construction

Each prototype pump (impeller and volute) was constructed using a rapid prototype

manufacturing process. The various configurations were drawn in CAD software

(SolidworksTM) and exported as STL files. These files had a chord angle of 3 degrees

and tolerance of 0.01mm. The STL files were sent to a rapid prototyping facility

(Concentric Asia Pacific, Brisbane, Australia) for manufacture using the stereo-

lithography process.


5-5

This process created each part from an epoxy based resin. Each piece was then post

processed by bead blasting and sanding all surfaces prior to assembly in the testing

facility. The testing facility, described in more detail later, was designed to

accommodate each modular pump component for subsequent hydraulic performance

and force testing.

5.2.2 Equipment / Instrumentation / Software / Code

Equipment / Instrumentation

JR3

Impeller hydraulic force data acquisition was achieved using the JR3 50M31 A-125

100N5 force transducer (JR3, California, USA). This force transducer has the ability

to resolve forces and moments in all six degrees of freedom at any one time. The

maximum measurable load in the axial (z) and radial (x,y) directions were 200N and

100N respectively. The maximum moment about all axes is 5Nm. The transducer

was manufactured 20/03/1997 and was recalibrated on 31/12/2003.

Pinch Valves

Proportional control valves (ECPV-375B, HASS Manufacturing, USA) were used to

control pump afterload resistance. The valves were powered with a 12VDC/2A

supply and proportionally responded to a 5VDC signal produced from two dSPACE

card DAC channels. The valves had the additional ability to alter actuation rate by

altering onboard electronic jumper configuration.

The jumpers were configured at the slowest rate of actuation (approx 30seconds open

to closed). Furthermore, the rate of control input voltage signal was set by the

SIMULINK model to reduce the actuation time to approx 120 seconds. This code

was implemented to automate the capture of non-pulsatile performance

characteristics.

DSPACE

The mock loop and assist device were simultaneously controlled by a digital signal

processing (DSP) controller board. The card was a 250 MHz model DS1104

(dSPACE, MI, USA) that incorporated a combined total of eight ADC (4x16bit,

4x12bit) and eight DAC (8x16bit) channels, 20 bit digital I/O, and two digital

incremental encoder interfaces (TTL compatible).


5-6

Motor

The VAD iterations were driven by a conventional DC motor AMAX32, Maxon

Motor, Switzerland). A 0-10V signal representing the set rotational speed was

determined from the SIMULINK speed controller and output from the dSPACE card.

This signal was amplified with a gain of 2.4 V/V to achieve the full 0 - 24 V range of

the motor. The actual motor speed was recorded with the integrated encoder, and fed

back to the speed controller via the dSPACE card.

Software / Code

MATLAB

Custom code was written in MATLAB (Jr3.m). The code interfaced with the JR3

transducer via the MATJR3.dll driver, developed by J. Norberto Pires. This enabled

the real-time capture of force and torque data for future analysis. The code and full

description of its operation is provided in APPENDIX D.

SIMULINK Model

A model was created in SIMULINK to provide a link between the experimental test

rig and the CONTROLDESK interface. This model replicated the components of the

experimental system, and incorporated parameters for real-time variation from

CONTROLDESK. A full description of the model is provided in APPENDIX D.

The model controls the operation of the mock circulation rig independent of VAD

status. The model also incorporates subsystems that dictate VAD operation.

VAD Subsystem A subsystem within the model records the performance and controls the operation of

the inserted VAD. This subsystem is responsible for recording and maintaining

rotational speed, recording motor power and determining pump efficiency.

CONTROLDESK

An interface program was developed in CONTROLDESK for the real-time capture

and manipulation of relevant parameters. Mock circulation loop operation was

determined, before pump rotational speed was decided and maintained. A splash

screen of the CONTROLDESK interface is provided in APPENDIX D.


5-7

5.2.3 Experimental Design

Evaluation of LVAD performance and impeller hydraulic force was conducted in-

vitro using the pulsatile mock circulation loop presented in Chapter 3. The loop

initially operated in a non-pulsatile configuration to identify pump performance

curves and non-pulsatile hydraulic forces over the entire flow capacity range. It was

then configured to replicate a degree of untreated left heart failure, before the LVAD

was introduced to assess its ability to re-establish hemodynamic parameters from

pathological to normal levels. The impact of failing heart function on hydraulic

forces was simultaneously evaluated.

Hydraulic Force and Performance Measurement

Each LVAD impeller was supported by a shaft, seal and contact bearing system.

Rotation was provided by a 20W motor and monitored with a digital encoder

(AMAX32 / HEDS 55__, Maxon Motor, Switzerland). The shaft, bearing and motor

assembly (grey) was separated from the pump casing (black) by a multidimensional

force transducer (50M31, JR3 Inc, USA) and two flexible couplings (Figure 5-2).

The hydraulic forces imposed on the impeller were transferred through the shaft and

resisted by the force transducer. The transducer exhibited high displacement

stiffness, preventing impeller movement within the pump cavity. To eliminate fluid

momentum effects on the shaft at the pump inlet eye, a metallic sleeve protruded

from the centre of the inlet volute.

Modular pump components enabled the interchangeable testing of all volute types for

each impeller type. Another single volute was created with ten pressure ports, which

were connected to pressure transducers for measuring volute pressures around the

impeller (Figure 5-3).


5-8

Figure 5-2 –Force Test Rig


5-9

0°

90°

180°

270°

5

4

3 21

7

6

10

8

9

Figure 5-3 – Pressure Tapping Locations

These pressure readings were integrated to reveal static radial forces at three flow

capacities (Shut-off, design, max flow). Performance and force data were

automatically gathered in real-time by interfacing signals to a PC using a dSPACE

controller board (DS1104, dSPACE, MI, USA) and JR3 inc. acquisition card. This

enabled the use of SIMULINK and CONTROLDESK for implementation of custom

control and acquisition code. Afterload vascular resistance was controlled with a

proportional control pinch valve (HASS Manufacturing, USA). Pressure and flow

rates were monitored with TOYODA pressure transducers (SD10B-1, Gambro, USA)

and 10mm ID electromagnetic flow meters (IFC010, KROHNE, Sweden).

5.2.4 Experimental Procedure

Non-Pulsatile Condition

Real-time performance and force data were recorded in the non-pulsatile mock loop

configuration at a constant pump speed while varying afterload resistance

Pump speed set in CONTROLDESK was maintained using a SIMULINK coded

speed controller, using feedback from the digital encoder. The proportional control


5-10

pinch valve was controlled to slowly open from shut-off, to automatically vary

afterload resistance while capturing pressure and flow signals in CONTROLDESK.

Radial (x,y) force, axial (z) force, and torque (Mz) data were simultaneously

recorded using custom MATLAB code interfacing to the JR3 force transducer via the

JR3 acquisition card. Impeller weight and dynamic unbalance were offset prior to

insertion in the mock loop.

Motor input power was recorded throughout the duration of each test, and divided by

the motor efficiency to reveal the total electrical input power (Pin = Volts x Amps /

Meff) supplied to each pump. The total hydraulic output power (Pout = pgHQ + c2/2g)

was derived from the recorded pressure and flow rate values. By dividing the output

power by the input power, a value of overall pump efficiency was revealed.

After completing initial tests, the effect on performance of axial impeller actuation

within the pumping cavity was investigated. The single volute, semi-open impeller

(112) LVAD configuration was inserted into the non-pulsatile loop, and operated at

2600 rpm with the impeller in the central position (middle). The circuit resistance

was then tuned to provide a constant 100mmHg afterload on the device. With the

use of 0.5mm spacers, the impeller was then moved axially relative to the casing

until the axial clearance above the vanes was minimal (top). Resistance was tuned

and complete performance curve results were recorded for the new position. Spacers

were used to increase the axial clearance in 0.5mm increments, each time recording

results, until a maximum clearance gap was observed (bottom). The pump speed and

afterload resistance were again controlled while the impeller was moved from bottom

to top location. The changes in flow rate were recorded in real-time. Finally, results

were also obtained for pump operation at 2400rpm at top, middle and bottom

impeller locations.

Pulsatile Heart Failure Condition

Failing heart function was initiated in the mock circulation loop, as described in

Chapter 3. The single volute, closed impeller type design was then introduced to

assess its ability to re-establish normal hemodynamic parameters. The impact of

failing heart function on hydraulic forces was simultaneously evaluated.


5-11

5.2.5 Calibration

Equipment and instrumentation was calibrated prior to testing. A brief description of

this process follows, with the results from the calibration tests provided in

APPENDIX E.

Electric Motor

To determine the actual input electrical power, the power required to overcome the

bearing resistances must be taken away from the total measured value. The pump

was run dry from 200 to 3000 rpm before each experiment, with values of Volts and

Amps recorded. The electrical power (V x A) was then multiplied by the maximum

motor efficiency (motor constant) (77%) to determine the power required to

overcome bearing resistance. The power was assumed to increase linearly with

respect to speed and therefore an equation was established that enabled the

appropriate no load power values to be subtracted from power readings recorded

from the pending experiments.

JR3 Force Transducer

The small range of recorded radial and axial hydraulic force presents a limitation due

to the transducer full scale range (100 N). The 14 bit resolution of 6 mN was

adequate for static force measurements, however, poor sensitivity meant a less than

ideal signal to noise ratio, preventing the investigation of high frequency dynamic

forces (up to 1 kHz). The influence of noise on the static force readings was removed

by digitally filtering the transducer signals to produce time averaged values.

To ascertain the actual values of force in each degree of freedom, unbalance forces

produced while operating the pump dry should be subtracted from the values

measured during wet operation. Therefore, the pump was run dry at speeds ranging

from 200 to 3000 rpm with values of force in each degree of freedom recorded.

Pressure Transducers

Each pressure transducer was calibrated before use. The voltage signal for

atmospheric pressure was recorded before a known pressure was exposed to each

transducer. The difference in pressure was divided by the change in voltage signal to

calculate the individual transducer gain (mmHg/Volt). The gains were then altered

to common values (770mmHg/V or 385mmHg/V) for simplicity.


5-12

(a) LVAD (b) RVAD

(c) Bi-VAD Side

Figure 5-4 – Experimental VAD configurations


5-13

5.2.6 Experimental Configurations

The LVAD (Figure 5-4(a)) configurations were inserted into the complete mock

circulation rig to assess performance characteristics and hydraulic forces. Not every

configuration was tested for force and performance data in non-pulsatile and pulsatile

modes. Table 5-1 cross-references the volute and impeller types, with the legend

indicating the tested conditions.

Table 5-1 – LVAD Impeller/Volute Experimental Configurations

Impeller Discharge Angle (β2)

Volute Type 22.5 112.5 Single Double Circular

P,N,F N,F N,F

N x x

P – Pulsatile Performance, N – Non-Pulsatile Performance, F - Forces Recorded, x - Not tested

A single RVAD (Figure 5-4(b)) was inserted into the loop to evaluate the design’s

hydraulic performance characteristics in the non-pulsatile loop only, for the

configurations described in Table 5-2. No impeller forces were recorded.

Table 5-2 – RVAD Impeller/Volute Experimental Configurations


Volute Type 22.5 112.5 Single Double Circular

N x N

x x x

N – Non-Pulsatile Performance, x - Not tested

The impact of Bi-VAD insertion (Figure 5-4(c)) on the hemodynamics of simulated

left and right heart failure (pulsatile) conditions was qualitatively assessed for

configuration B (Table 5-3). Again, no impeller forces were recorded.

Table 5-3 – BVAD Impeller/Volute Experimental Configurations

LEFT Right

Impeller Discharge Angle (b2) Impeller Discharge Angle (β2) Volute Type 22.5 122.5 22.5 112.5

Single Double Circular

(A) x x x

x x

(B) P

(A) x -

(B) P

x x x

(A) Pump Configuration 1, (B) Pump Configuration 2, P – Pulsatile Performance, x - Not tested

The current iteration Bi-LVAD was not tested, however the performance and force

data from a previous iteration was recorded and presented in (Timms, Tan et al.

2004).


5-14

0 2 4 6 8 100

50

100

150

200

Flow Rate (L/min)

Pres

sure

(mm

Hg)

Single Volute Performance Characteristics (β22.5°)

⊕

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

0 2 4 6 8 100

10

20

30

40

Flow Rate (L/min)

Effic

ienc

y (%

)

Single Volute Efficiency (β22.5°)

2700Rpm2400Rpm2100Rpm1800Rpm

(Single Volute, 22.5o)

0 2 4 6 8 100

50

100

150

200

Flow Rate (L/min)

Pres

sure

(mm

Hg)


⊕

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

0 2 4 6 8 100

10

20

30

40

Flow Rate (L/min)

Effic

ienc

y (%

)



(Single Volute, 112.5o)

0 2 4 6 8 100

50

100

150

200

Flow Rate (L/min)

Pres

sure

(mm

Hg)

Double Volute Performance Characteristics

⊕

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

0 2 4 6 8 100

10

20

30

40

Flow Rate (L/min)

Effic

ienc

y (%

)

Double Volute Efficiency


(Double Volute, 22.5o)

0 2 4 6 8 100

50

100

150

200

Flow Rate (L/min)

Pres

sure

(mm

Hg)

Circular Volute Performance Characteristics

⊕

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

0 2 4 6 8 100

10

20

30

40

Flow Rate (L/min)

Effic

ienc

y (%

)

Circular Volute Efficiency


(Circular Volute, 22.5o)

Figure 5-5 – LVAD Family of Pump Performance and Efficiency Curves


5-15

5.3 Results

5.3.1 Non-Pulsatile

Hydraulic Performance

LVAD

Performance Characteristics The pump characteristic and stage efficiency curves for each of the LVAD

configurations inserted into the non-pulsatile loop are displayed in Figure 5-5. The

rotational speed required to achieve design conditions depended on the

impeller/volute configuration (Table 5-4).

Table 5-4 - Design Point Rotational Speed (RPM) of LVAD Impeller/Volute Configurations

Impeller Discharge Angle (β2) Volute Type 22.5o 112.5o

Single Double Circular

2400 RPM 2475 RPM 2450 RPM

2550 RPM x x

The single volute (22.5) configuration (where the bracketed number corresponds to

the discharge angle) featured the lowest rotational speed and consequently the

highest “best efficiency”. A slight increase in rotational speed and corresponding

decrease in maximum efficiency was observed for the double volute (22.5), circular

(22.5) and single volute (112.5) configurations respectively.

All pump curves were relatively flat below design point. All curves, except the single

volute (112), tapered down significantly above the design point. Instead, the single

volute (112) configuration maintained a relatively flat response, as anticipated for

pumps employing forward facing vanes.

The double volute (22.5) configuration could not attain the same level of best

efficiency as the single volute (22.5). However efficiency’s at off design flow rates

were higher. Single volute (112) performance exhibited the least efficiency at design

conditions.

Results for the single volute (90) configuration are not presented due to poor

performance. This was attributed to excessive shroud blockage due to a small

impeller vane height, and will be discussed in further detail later in the chapter.


5-16

Variation of Axial Clearance The effect on performance of axial clearance variation (between the vanes and top

casing) of the semi-open (112) impeller is presented in Figure 5-6. It is observed

that efficiency and thus hydraulic performance over all capacities is best when axial

clearance is minimised (0.1mm). As clearance is then increased, a corresponding

decrease in efficiency and performance is observed until the clearance is maximised

(2.5mm).

Maintaining pump speed (2600 rpm) and afterload resistance (100mmHg) while

actuating the impeller axially, produced flow rates and efficiencies contained within

the limits specified by the performance characteristic curves.

Flow rate was therefore observed to range between 4.8 – 6.1 L/min despite the pump

operating against a constant afterload and rotational speed.

0 2 4 6 8 100

50

100

150

200

Flow Rate (L/min)

Pres

sure

(mm

Hg)

2600 RPM Performance (Axial Clearance 0.1-2.5mm)

⊕

0.1mm0.5mm1.0mm1.5mm2.0mm2.5mm

0 2 4 6 8 100

5

10

15

20

25

30

35

40

Flow Rate (L/min)

Effic

ienc

y (%

)

2600 RPM Efficiency (Axial Clearance 0.1-2.5mm)

0.1mm0.5mm1.0mm1.5mm2.0mm2.5mm

(a) Performance (left) and efficiency (right) at constant speed and variable afterload

0 2 4 6 8 100

50

100

150

200

Flow Rate (L/min)

Pres

sure

(mm

Hg)


2600 Rpm

0 2 4 6 8 100

10

20

30

40

Flow Rate (L/min)

Effic

ienc

y (%

)


2600Rpm

(b) Performance (left) and efficiency (right) at constant 2600rpm and 100mmHg afterload

Figure 5-6 – LVAD Performance with Constant Speed and Axial Clearance Variation


5-17

Non-Dimensional Parameters Specific head (hs) was plotted against specific capacity (qs) for direct comparison of

each configuration’s hydraulic performance (Figure 5-7). This graph is non-

dimensional, therefore each family of performance curves are grouped into a single

non-dimensional curve.

0 0.005 0.01 0.015 0.02 0.025 0.03 0.0350

0.1

0.2

0.3

0.4

0.5

0.6

0.7

0.8

0.9

1

Flow Coefficient qs

Hea

d Co

effic

ient

ψ

Non-Dimensional Performance Characteristics

Single Volute (β22.5°)Single Volute (β112.5°)Double Volute (β22.5°)Circular Volute (β22.5°)

Figure 5-7 – LVAD Non-Dimensional Pump Performance Curves

Maximum capacity is achieved in both single volute (22.5 & 112.5) configurations,

with the latter exhibiting a flatter characteristic curve, albeit at lower specific head.

Both circular and double volute maximum specific capacity is limited, while the

former exhibits greatest specific head at shutoff. The specific head at shut-off of

each backward blade configuration ranged between 0.5 and 0.6.

Specific Power was also calculated and plotted against specific capacity. This non-

dimensional graph (Figure 5-8) is used to compare the power requirements of the

various pump configurations. Due to difficulties in obtaining exact input power

measurements, the graph is presented to outline the qualitative effect of power trends

for each volute configuration. For instance, all configurations incorporating


5-18

backward swept discharge blade angles exhibited a decreasing power gradient with

increasing capacity. The forward facing blade angle configuration however,

produced a slightly increasing power gradient. Furthermore, although the power

value accuracy is questionable, the results indicate the presence of a power

requirement due to recirculation at shut-off conditions.

0 0.005 0.01 0.015 0.02 0.025 0.03 0.0350.02

0.025

0.03

0.035

0.04

0.045

0.05

Specific Capacity (qs)

Spec

ific

Pow

er (P

s)

Specific Power (Ps) Vs Specific Capacity (qs)

Single Volute (β22.5°)Single Volute (β112.5°)Double Volute (β22.5°)Circular Volute (β22.5°)

Figure 5-8 – LVAD Non-Dimensional Power Curves


5-19

RVAD

Performance Characteristics The family of performance curves and non-dimensional graph for each RVAD

configuration are presented in Figure 5-9 and Figure 5-10 respectively.

0 1 2 3 4 5 60

5

10

15

20

25

30

35

40

Flow Rate (L/min)

Pres

sure

(mm

Hg)

Single Volute Performance Characteristics

⊕

3000 Rpm2700 Rpm2400 Rpm2100 Rpm1800 Rpm

0 1 2 3 4 5 60

5

10

15

20

25

30

35

40

Flow Rate (L/min)

Pres

sure

(mm

Hg)

Circular Volute Performance Characteristics

⊕

3000 Rpm2400 Rpm2100 Rpm1800 Rpm

(Single Volute, 22.5o) (Circular Volute, 22.5o)

Figure 5-9 – RVAD Family of Pump Performance Curves

0 0.05 0.1 0.15 0.2 0.25 0.30

0.1

0.2

0.3

0.4

0.5

0.6

Specific Capacity φ

Spec

ific

Hea

d ψ

Specific Head (hs) Vs Specific Capacity (qs)

Single Volute (β22.5°)Circular Volute (β22.5°)

Figure 5-10 – RVAD Non-dimensional Pump Performance Curves


5-20

The results characterise the poor performance of each configuration, with the design

condition met at rotational speeds well above the theoretical speed of 1921rpm

(Table 5-5).

Table 5-5 - Design Point Rotational Speed (RPM) of RVAD Impeller/Volute Configurations


Volute Type 22.5 Single

Circular 3000 RPM 3800 RPM

The non-dimensional analysis also confirms this performance, with a common

maximum specific head at shut-off of 0.4, and maximum flow coefficients of 0.14

and 0.19 for the circular and single volutes respectively.

This poor performance was attributed to flow blockage by the 6 impeller vanes.

Furthermore, the inlet swirl effect of the inlet volute was not considered, and could

contribute to a reduction in useful output head.

Considerable scattering was observed in the characteristic RVAD curves, which is

not consistent with conventional pump operation. This scattering could be a result of

the automated method of pressure and flow recording. The variation of outlet

resistance was not smooth, as the pinch valve steps its position from close to open.

After each step, pressure immediately drops while the inertiance of the flow

contributes to the slower response to increase flow rate. Since both pressure and flow

are recorded in real-time, these variations are captured as scatter.


5-21

Hydraulic Force

Static Radial Force

Static radial force magnitudes and directions are displayed in Figure 5-11 for each

volute type.

0 2 4 6 8 100

0.2

0.4

0.6

Flow Rate (L/min)

Forc

e (N

)

Single Volute Radial Thrust Magnitude

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

0.25

0.5

30

210

60

240

90

270

120

300

150

330

180 0

Single Volute Radial Thrust Direction

Force X (N)

Forc

e Y

(N) Q <Qn

Q =QnQ>Qn

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

(a)i. Single Volute Magnitude (a)ii. Single Volute Direction

0 2 4 6 80

0.2

0.4

0.6

Flow Rate (L/min)

Forc

e (N

)

Double Volute Radial Thrust Magnitude

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

0.2

0.4

30

210

60

240

90

270

120

300

150

330

180 0

Double Volute Radial Thrust Direction

Force X (N)

Forc

e Y

(N)

Q = 0

Q = Qn

Q > Qn

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

(b)i. Double Volute Magnitude (b)ii. Double Volute Direction

0 2 4 6 8 100

0.2

0.4

0.6

Flow Rate (L/min)

Forc

e (N

)

Circular Volute Radial Thrust Magnititude

2400 Rpm2100 Rpm1800 Rpm

0.5

1

30

210

60

240

90

270

120

300

150

330

180 0

Circular Volute Radial Thrust Direction

Force X (N)

Forc

e Y

(N)

Q<Q n Q=QnQ>Qn

2400 Rpm2100 Rpm1800 Rpm

(c)i. Circular Volute Magnitude (c)ii. Circular Volute Direction

Figure 5-11 – Radial Thrust in single (a), double (b) and circular (c) volutes.

0°

90°

180°

270°

0

90

180

270

0

90

180


5-22

Table 5-6 compares the radial thrust magnitudes at the design point

Table 5-6 – Radial trust at design point (5L/min @ 100mmHg)

Volute Radial Force (N) Single Double Circular

0.01 0.05 0.5

The trend of radial thrust versus non-pulsatile flow capacity follows that described

by theory. The single volute (Figure 5-11(a)i) experienced the smallest radial thrust

at the design point, followed by the double (Figure 5-11 (b)i) and circular volute

(Figure 5-11(c)i). The direction of the force vector depended on the pump

operating point in relation to the design point. For example, the force acted away

from the cutwater at below design conditions in the single volute, while it pointed

toward the cutwater above design (Figure 5-11(a)ii). This is explained by the degree

of fluid acceleration and thus velocity upstream of the cutwater. A similar trend is

found in the double volute, though at smaller magnitudes (Figure 5-11(b)ii). Radial

thrust was directed toward the cutwater of the circular volute at all capacities (Figure

5-11(c)ii).

Pressure Tappings Integration of volute pressure was used to verify the transducer results. Operation of

the pump at 2400 rpm revealed a maximum force of 0.32N at shut-off, 0.05N at

design, and 0.27N at maximum capacity (Figure 5-12).

0.1

0.2

0.3

0.4

30

210

60

240

90

270

120

300

150

330

180 0

Radial Force Direction and Magnitude at Various Operating Conditions

Force X (N)

Forc

e Y

(N) Q = 0

Q = QnQ > Qn

Figure 5-12 – Radial Thrust Direction & Magnitude Integrated from Single Volute Pressures

Readings at Shut-off, Design and Maximum Flow Capacities.


5-23

Static Axial Force

The static axial forces encountered by the closed and open type impellers over all

non-pulsatile flow rates are presented in Figure 5-13.

0 2 4 6 8 100

2

4

6

8

Flow Rate (L/min)

Axi

al F

orce

(N)

Closed Impeller Axial Thrust Magnitude

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

0 2 4 6 8 10

0

2

4

6

8

Flow Rate (L/min)

Axi

al F

orce

(N)

Semi-Open Impeller Axial Thrust Magnitude

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

(a) Closed Impeller (b) Semi-Open Impeller

Figure 5-13 – Axial Thrust on Impeller Types

Table 5-7 compares values of axial thrust at the design point. As expected, axial

thrust of the semi-open type impeller was larger in magnitude than the closed type

impeller.

Table 5-7 – Axial thrust at design point

The technique of axial force balancing using a double entry impeller was evaluated

during the investigation of a previous iteration (Timms, Tan et al. 2004). Impeller

and pump geometry was therefore dissimilar to the current iterations presented.

However, the results indicated a reduction in axial force at design point from 7.55N

to 0.84N for single and double sided semi open impellers respectively (Figure 5-14).

Double Pump Axial Force

Figure 5-14 – Comparison of Axial Thrust in a Semi-Open single and double impeller

(Timms, Tan et al. 2004)

Impeller Type Axial Force (N) Closed Semi-open

3 4.5

halla



5-24

Reaction Torque

Values of reaction torque are displayed in Figure 5-15. The values are not absolute,

since an unknown transducer offset was eliminated by returning the value of torque

at shutoff capacity to zero. This is not entirely practical, as some torque and thus

power is required to recirculate the fluid at zero capacity. Instead, the result describes

the relative magnitudes of torque encountered at each rotational speed for the single

volute configuration. The resulting torque of 0.0075 Nm at design point is slightly

lower, but in the same order to that expected of a pump operating at the efficiencies

outlined previously.

0 2 4 6 8 10-0.02

-0.015

-0.01

-0.005

0

Flow Rate (L/min)

Reaction Torque

2700 Rpm2400 Rpm2100 Rpm1800 Rpm

Figure 5-15 – Reaction Torque in Single Volute (22.5)

Comparison to Theory

The hydraulic performance characteristic results (pressure and flow) were used in the

equations proposed by each author (Appendix C) to predict radial and axial thrust

characteristics. Comparisons were then made to the recorded thrust values to

determine the most appropriate equation for thrust prediction in each impeller/volute

configuration.

Magnitude

Radial Thrust

Non-dimensional “radial thrust factors” for each volute type were calculated from

recorded values of radial thrust. These were compared to predictions made by each

author for similar pump operations (Figure 5-16).


5-25

0 20 40 60 80 100 1200

0.05

0.1

0.15

0.2

0.25

0.3

0.35

0.4

Percent Capacity of B.E.P (%)

Non

-Dim

ensio

nal R

adia

l Thr

ust F

acto

r (K

)

Comparison of Practical to Theoretical Radial Thrust Factors (K)

Single VoluteDouble VoluteCircular VoluteTheory SV (KARASSIK)Theory DV (ANSI)Theory CV (STEPANOFF)

Figure 5-16 – Comparison of measured Radial thrust with theoretical prediction

Good agreement was found between experimental single volute radial thrust factor

and that predicted by Karassik (2000). Estimation of double volute results was best

made using the method proposed in the ANSI (ANSI/HI 1994). Stepanoff’s method

of calculating radial thrust (Stepanoff 1957) provided the most accurate comparison

to experimental circular volute results.

Axial Thrust

Axial thrust in the closed type impeller was found to be relatively independent of

volute type, aside from the influence on pressure development. Axial thrust

predictions made by Stepanoff (1957) described a similar trend to the experimented

results, however at a slightly lower magnitude (Figure 5-17).


5-26

0 20 40 60 80 100 1200

1

2

3

4

5

Percent Capacity of B.E.P (%)

Axi

al F

orce

(N)

Comparison of Practical to Theoretical Axial Thrust

Single VoluteDouble VoluteCircular VoluteAll Volutes (Stepanoff)

Figure 5-17 – Comparison of measured axial thrust with theoretical prediction

Direction The directions of radial thrust experienced by the impeller operating from shut-off to

full capacity at 2400 rpm in the single, double and circular volutes are presented for

comparison in Figure 5-18. This figure illustrates the dependence of thrust direction

on the volute type employed.

0°

90°

180°

270° Single

0

90

180

270 Double

0

90

180

0.5

1

30

210

60

240

90

270

120

300

150

330

180 0

Comparison of Radial Thrust Direction (2400 rpm)

Force X (N)

Forc

e Y

(N)

Circular VoluteDouble VoluteSingle Volute

Circular

Figure 5-18 – Direction of Radial Thrust in Single, Double and Circular Volutes at 2400rpm

Q<Qn

Q>Qn Q<Qn

Q>Qn

Q=0


5-27

The single volute result was then compared to predictions of direction made by

(Agostinelli, Nobles et al. 1960) in Figure 5-19 at similar specific speeds to the

pump used in this study (Ns = 740). The original figure (Figure 4-13) was re-oriented

to match the outlet geometry of the pump for the experimental results. That is, the

pump outlet directions were arranged to coincide with the same axis. Reasonable

correlation was found, with the experimental shut-off and full capacity directions

located in the predicted quadrants.

As expected, the direction of axial thrust was always toward the impeller inlet eye

(+Z), due to the imbalance of pressures acting on the impeller top and underside.

0

90

180

270

Ns=530

Ns=784

Q<Qn

Q>Qn

Figure 5-19 – Radial thrust direction of comparative specific speeds in a single volute


5-28

0.8 1 1.2 1.4 1.6 1.80

20

40

60

80

100

120

Time (sec)

Pres

sure

(mm

Hg)

Untreated Heart Failure (Systemic)

LAPLVPAoPMAP

0.8 1 1.2 1.4 1.6 1.8

0

20

40

60

80

100

120

Time (sec)

Pres

sure

(mm

Hg)

LVAD Support (Systemic)

LAPLVPAoPMAP

(i) Systemic Pressure Distribution (i) Systemic Pressure Distribution

0.8 1 1.2 1.4 1.6 1.80

5

10

15

20

25

30

35

40

Time (sec)

Pres

sure

(mm

Hg)

Untreated Heart Failure (Pulmonary)

RAPRVPPAPMPAP

0.8 1 1.2 1.4 1.6 1.8

0

5

10

15

20

25

30

35

40

Time (sec)

Pres

sure

(mm

Hg)

LVAD Support (Pulmonary)

RAPRVPPAPMPAP

(ii) Pulmonary Pressure Distribution (ii) Pulmonary Pressure Distribution

Figure 5-20 – Untreated heart failure pressure for systemic (i) and pulmonary (ii) circulation.

Figure 5-21 – LVAD supported pressure for systemic (i) and pulmonary (ii) circulation.

0 10 20 30 40 50 60 70 80 90 1002

2.5

3

3.5

4

4.5

5

5.5

Time (sec)

Perf

usio

n Ra

te (L

/min

)

Heart Failure Transition to LVAD Support (2200 rpm)

⊕ LVAD Insertion Point

⊕LVAD Support MSQ

SQMPQPQ

Figure 5-22 – Transition of perfusion from Heart Failure to LVAD Support


5-29

5.3.2 Pulsatile

Left Heart Failure with LVAD Support


Individual systemic and pulmonary pressures and flow rates were measured in the

mock circulation loop during heart failure, and subsequent single volute LVAD

support, to determine the hemodynamic effect of VAD insertion.

Pressures Simulated pathological blood pressures experienced in the heart failure condition

(Figure 5-20) were returned to acceptable levels with LVAD support (Figure 5-21)

in the (i) systemic and (ii) pulmonary circulatory systems. Detailed pressure values

are provided in Table 5-8. In particular, left atrial pressures were reduced from 25 to

8 mmHg, indicating the minimisation of pulmonary congestion. Mean aortic pressure

(MAP) was also improved. Left ventricular pressure (LVP) did not exceed the MAP

value during systole, suggesting the aortic valve remained closed. Finally, the aortic

pulse pressure was reduced, but not eliminated.

Table 5-8 – Comparison of Heart Failure to LVAD Support Component Pressures

mmHg Heart Failure LVAD LAP 25 8 LVP 100 60

LVPED 25 5 AoP 95/70 99/89 MAP 82 95 RAP 20 3 RVP 40 33

RVPED 17 3 PAP 35/28 28/11

MPAP 31 18 Pmc 20 7

Perfusion Successful cardiac output was re-established from pathological (2.3L/min) to

acceptable (5.1 L/min) perfusion levels while operating the pump at a constant

2200rpm (Figure 5-22).


5-30

0 1 2 3 4 5 6 7 80

20

40

60

80

100

120

140

160

Flow Rate (L/min)

Pres

sure

(mm

Hg)

LVAD Performance Characteristics for the Support Condition

⊗

⊕

Diastole

Systole

2400 Rpm2200 Rpm2100 Rpm

Figure 5-23 – Non-Pulsatile LVAD operation in the pulsatile LHF environment

0 1 2 3 4 5 6 7 8 9 100

5

10

15

20

25

30

35

40

Flow Rate (L/min)

Effic

ienc

y (%

)

LVAD Efficiency for the Support Condition

⊗

⊕

Diastole

Systole

2400 Rpm2200 Rpm2100 Rpm

Figure 5-24 – Non-Pulsatile LVAD efficiency in the pulsatile LHF environment


5-31


LVAD performance during the pulsatile support condition is displayed in Figure

5-23, where the pump operates along the interpolated characteristic curve for 2200

rpm. The limits on this curve correspond to the pressure difference from the left

ventricle to the aorta throughout the entire cardiac cycle, i.e. from systolic

(35mmHg) to diastolic (90mmHg) periods. The average flow rate for each cardiac

cycle (5.1 [L/min]) is determined from the mean time spent in systole (40%) and

diastole (60%) (Table 5-9).

Table 5-9 – Pressure and flow conditions for diastolic and systolic periods

Diastole Systole Pressure (mmHg) Pump ∆P 90 35 LVP 5 60 AoP 95 95 Perfusion (L/min) 4.2 6.5

The pump stage efficiency curve was interpolated and displayed in Figure 5-24.

Efficiency rose from a diastolic value of 30 % to briefly peak at 33%, before falling

to 23% during systole.


5-32

25 30 35 400

0.05

0.1

0.15

0.2

0.25

0.3

Diastole

Systole

Time (Sec)

Radi

al F

orce

(N)

LVAD Support Radial Force Magnitude

0 1 2 3 4 5 6 70

0.05

0.1

0.15

0.2

0.25

0.3

⊕

⊗

Systole

Diastole

Flow Rate (L/min)

Resu

ltant

Rad

ial F

orce

(N)

Comparison of Radial Force

2400 Rpm2200 Rpm2100 Rpm

(a) Radial Magnitude (a) Radial Magnitude

0.1

0.2

0.3

30

210

60

240

90

270

120

300

150

330

180 0

LVAD Support Radial Force Direction

Force X (N)

Forc

e Y

(N)

Systole

Diastole

0.2

0.4

30

210

60

240

90

270

120

300

150

330

180 0

Comparison of Radial Force Direction

Force X (N)

Forc

e Y

(N)

Systole

Diastole

⊕

⊗

2400 Rpm2100 Rpm

(b) Radial Direction (b) Radial Direction

25 30 35 400

0.5

1

1.5

2

2.5

3

3.5

Diastole

Systole

Time (Sec)

Axi

al F

orce

(N)

LVAD Support Axial Force Magnitude

0 1 2 3 4 5 6 70

0.5

1

1.5

2

2.5

3

3.5

Systole

Diastole

⊕

⊗

Flow Rate (L/min)

Axi

al F

orce

(N)

Comparison of Axial Force

2400 Rpm2200 Rpm2100 Rpm

(c) Axial Magnitude (c) Axial Magnitude

25 30 35 40-0.01

-0.008

-0.006

-0.004

-0.002

0

Systole

Diastole

Time (Sec)

Torq

ue (N

m)

LVAD Support Reaction Torque

0 1 2 3 4 5 6 7-0.01

-0.008

-0.006

-0.004

-0.002

0

SystoleDiastole

⊕⊗

Flow Rate (L/min)

Torq

ue (N

m)

Comparison of Reaction Torque

2400 Rpm2200 Rpm

(d) Reaction Torque (d) Reaction Torque

Figure 5-25 – Fluctuation of force with systole and diastole

Figure 5-26 –Forces from pulsatile and non-pulsatile test results

0°

90°

180°

270°

0°

90°

180°

270°


5-33

Hydraulic Force

Impeller radial force magnitude (Figure 5-25(a)) and direction (Figure 5-25(b)),

axial hydraulic force magnitude (Figure 5-25(c)) and motor reaction torques (Figure

5-25(d)), were recorded in the LVAD supported simulated heart failure environment.

The force magnitudes fluctuated within each simulated heat beat. Values

corresponding to diastolic and systolic periods are compared in Table 5-10. Radial

force direction changed for systolic and diastolic periods, while axial force was

always directed toward the pump inlet eye.

Table 5-10 – LVAD Support Force/Torque Magnitudes at Systole and Diastole

The recorded LVAD support forces were cross-referenced with previous non-

pulsatile values at corresponding flow rates for systole and diastole. The magnitudes

and directions were found to replicate interpolated non-pulsatile test results for 2200

rpm (Figure 5-26).

Radial force magnitude reciprocated the path outlined in Figure 5-26(a) for each

cardiac cycle. The force reduced to the minimum design point value during each

transition, resulting in a maximum change of 0.18N. The direction of the radial force

vector (Figure 5-26(b)) pointed away from the cutwater during diastole (110o), and

reversed direction to point toward the final volute section (310o) during the systolic

period. Figure 5-26(c) describes the 0.6N fluctuation of axial force magnitude

between systolic and diastolic limits corresponding to interpolated non-pulsatile axial

force data. Finally, a minimal 1.2mNm change in reaction torque was observed in

Figure 5-26(d) for this level of LVAD support.

Force Systole Diastole

Radial (N) 0.21 0.11

Axial (N) 2.0 2.6

Toque (mNm) 7.1 5.9


5-34

0 1 2 3 41618202224

Time (sec)

Pre

ssur

e (m

mH

g)


LAP

0 1 2 3 40

50

100

Time (sec)

Pre

ssur

e (m

mH

g)


LVP

0 1 2 3 4

50

60

70

Time (sec)

Pre

ssur

e (m

mH

g)(d) Aortic Pressure

AoPMAP

2.4 2.6 2.8 3 3.2 3.40

20

40

60

80

100

120

Time (sec)

Pre

ssur

e (m

mH


LAPLVPAoPMAP

0 1 2 3 48

1012141618

Time (sec)

Pre

ssur

e (m

mH

g)


LAP

0 1 2 3 40

50

100

Time (sec)

Pre

ssur

e (m

mH

g)


LVP

0 1 2 3 480

90

100

Time (sec)

Pre

ssur

e (m

mH

g)

(d) Aortic Pressure

AoPMAP

1.2 1.4 1.6 1.8 2 2.20

20

40

60

80

100

120

Time (sec)

Pre

ssur

e (m

mH

g)


LAPLVPAoPMAP


1 2 3 40

20

Time (sec)

Pre

ssur

e (m

mH

g)


RAP

0 1 2 3 40

20

40

Time (sec)

Pre

ssur

e (m

mH

g)


RVP

0 1 2 3 420

30

40

Time (sec)

Pre

ssur

e (m

mH

g)


PAPMPAP

2.4 2.6 2.8 3 3.2 3.40

5

10

15

20

25

30

35

40

Time (sec)

Pre

ssur

e (m

mH

g)


RAPRVPPAPMPAP

0 1 2 3 40

20

Time (sec)

Pre

ssur

e (m

mH

g)


RAP

0 1 2 3 40

20

40

Time (sec)

Pre

ssur

e (m

mH

g)


RVP

0 1 2 3 420

30

40

Time (sec)P

ress

ure

(mm

Hg)


PAPMPAP

1.2 1.4 1.6 1.8 2 2.20

5

10

15

20

25

30

35

40

Time (sec)

Pre

ssur

e (m

mH

g)


RAPRVPPAPMPAP


Figure 5-27 – Systemic (a) and Pulmonic (b) BI-HF pressure distribution

Figure 5-28 – Systemic (a) and Pulmonic (b) BI-VAD supported pressure distribution

0 0.5 1 1.5 2 2.5 3 3.5 40

1

2

3

4

5

6

Time (sec)

Per

fusi

on R

ate

(L/m

in)

Perfusion Rate

MSQSQMPQPQ

0 0.5 1 1.5 2 2.5 3 3.5 4

0

1

2

3

4

5

6

Time (sec)

Per

fusi

on R

ate

(L/m

in)

Perfusion Rate

MSQSQMPQPQ

Figure 5-29 – Perfusion Rate for Rest Figure 5-30 – Perfusion Rate for Left Heart Failure.


5-35

Heart Failure with Bi-VAD Support

Bi-VAD configuration (B) was briefly tested under conditions of heart failure. The

right impeller incorporated a 22.5 degree backward impeller with no shroud cased in

a single volute, while the left impeller had a 112.5 degree forward impeller with no

shroud in a circular casing.


Figure 5-27 to Figure 5-30 describes the hemodynamic effect of Bi-VAD insertion

into a simulated heart failure condition. With appropriate manipulation of systemic

and pulmonary vascular resistances, physiological conditions representative of

normal heart function where re-established.

Mean Aortic pressure was improved from 60 to 93mmHg, although mean pulmonary

artery pressure was only slightly reduced from 30 to 28mmHg. A reduction in left

atrial pressure (20 to 8mmHg) was also observed.

The high level of pulmonary arterial pressure is due to the remaining right heart

function. That is, it would appear that the right heart function was not sufficiently

reduced to correctly imitate biventricular failure. Furthermore, leakage from left to

right sides of the pump also contributes to the high pulmonary pressure. In any case,

techniques described in later chapters may aid in the reduction of the contribution of

pulmonary pressure by the VAD.

Bi-VAD support established an acceptable perfusion level of 5.4 L/min from the 2.7

L/min encountered during simulated heart failure.

Although the results indicate the somewhat successful preliminary operation of the

Bi-VAD in re-establishing perfusion, the results would benefit from future

optimisations.


5-36

5.4 Discussion

5.4.1 Experimental Method

Isolation of impeller hydraulic force was successfully achieved by separating the

impeller, bearing and motor drive assembly from each volute casing. Mounting this

assembly on the high stiffness force transducer prevented impeller movement within

the pump cavity while accurately recording force and torque magnitudes in each

degree of freedom. The flexible seals prevented transmission of force from the

impeller assembly to the volute casing. The installation of the small metallic sleeve

around the shaft at the inlet eye was effective in eliminating the momentum effect of

the fluid entering the pump. Without such sleeve, the fluid jet streams would act on

the shaft, thus creating an additional radial force, albeit low in magnitude.

Any impeller dynamic instability was eliminated by acquiring offset values during

dry pump operation. The influence of noise on the static force readings in the non-

pulsatile environment was removed by digitally filtering the transducer signals to

produce time averaged values. The level of filtering did not impose on the magnitude

of force variation at the heartbeat frequency during pulsatile heart failure simulation.

The 100N full scale range of the JR3 force transducer presented a limitation with

respect to resolution, sensitivity and noise, preventing the investigation of high

frequency dynamic forces. Furthermore, transient pressure surges due to mock loop

valve closure also masked dynamic results. Although advanced signal processing

techniques could be applied to the raw data to improve dynamic results, these high

frequency dynamic force magnitudes are not expected to dominate the static resultant

force data.


5-37

5.4.2 Non-Pulsatile Operation

The simultaneous capture of non-pulsatile hydraulic performance and force data was

successfully conducted while maintaining pump speed and setting the actuation rate

of the afterload resistance valve from shut-off. This automated technique provided a

quick and efficient means of testing the multiple impeller/volute configurations.


Performance characteristic curves and non-dimensional data are discussed in the

following sections for the LVAD and RVAD impeller/volute configurations.

LVAD

Performance Characteristics The LVAD design point of 5L/min @ 100mmHg was achieved at 25% higher

rotational speed (2400rpm) than theoretically designed (1921rpm), indicating an

efficiency loss.

The family of performance curves (Figure 5-5) indicate pump characteristics at

design and off design capacities. The non-dimensional performance graph of specific

head versus specific capacity (Figure 5-7) groups these results and compares the

characteristics of each impeller volute configuration. This information is useful in

determining the response of the device to changes in physiological conditions,

induced by vascular resistance alteration. A flat performance curve is desirable, as

drastic changes in rotational speed are not required to increase perfusion while

maintaining sufficient mean arterial pressure. A flatter performance curve is

characteristic of impellers incorporating high discharge blade angles. The (112.5)

impeller displayed the flattest response, particularly above design, where pressure

fell considerably in this region for each of the low discharge angled impellers.

The non-dimensional shut-off head range of 0.5 - 0.6 for each backward facing blade

configuration is consistent with that identified by Smith (2004) in a study of blood

pump non-dimensional data. The shut-off head of 0.4 for the (112.5) impeller

indicates a relatively larger amount of slip and separation within the impeller blades.


5-38

Efficiency Efficiencies are also compared in Figure 5-5 for each impeller/volute configuration.

The single volute with backward discharge blade angle (22.5o) LVAD configuration

exhibited the best efficiency at the design point. The smooth transition of flow from

impeller to volute contributes to this effect, since volute velocity is constant around

the impeller periphery at this point.

The relative drop in best efficiency of the double volute configuration is attributed to

the resistive effect of the splitter plate and additional separation loss contributed by

the second cutwater. However, efficiencies at off design conditions (particularly part

flow) are significantly higher than the single volute. This is due to the effect of the

impeller discharging fluid into a more uniform pressure distribution (Stepanoff

1957). The splitter also restricts the highest achievable capacity, as the extra

boundary layers grow in size and contribute to blockage of the flow passage. This

effect would be even more pronounced in higher viscosity fluids (blood) than the

fluid (water) used in these experiments.

The reduction in circular volute configuration efficiency is attributed to the

separation and formation of eddy currents at cutwater. The turbulent fluid transition

from volute to throat generates significant eddy currents and recirculation, which

contributes to losses and effectively partially blocks the throat area. This

recirculation is observed in the previous flow visualisation study (Tan, Timms et al.

2004).

The lower efficiency of the 112.5o impeller discharge angle in the single volute

configuration is expected in pumps employing forward facing blades, and is due to

the inability of the volute to efficiently convert the high speed fluid jet streams off

the impeller into pressure.

The specific power results displayed in Figure 5-8 provide an analysis of the power

requirements of each impeller/volute configuration with increasing capacity.

Difficulties in obtaining exact measurements of input shaft power contribute to offset

variations between volutes, thus the accuracy of the values are questionable.

However, a qualitative trend of specific power with capacity is evident. As expected,

the backward swept impellers displayed a reducing gradient of power with capacity,

while the forward blades characteristically observed a slight increase in gradient.


5-39

Axial Clearance Variation The effect on hydraulic performance and efficiency of semi-open impeller axial

clearance variation in a variable afterload environment at constant rotational speed

was investigated and displayed in Figure 5-6(a). By progressively increasing the

clearance gap, more fluid would escape over the impeller vanes, thus reducing the

hydraulic efficiency and consequently pressure and flow.

The reduction in efficiency as the gap increased from 0.1 - 2.5mm translated to a

reduction in output flow rate while a constant afterload pressure was maintained.

Figure 5-6(b) displays this variation in flow (4.8 - 6 L/min) and efficiency (10 -

21%) for a constant afterload of 100mmHg.

This effect is dependant on the ratio of blade height to clearance gap. Therefore, the

change in efficiency for a given impeller axial displacement is more pronounced in

impellers with small vane heights. The type of impeller also contributes to the

magnitude, with semi-open impellers allowing for the greatest variation in efficiency.

RVAD

The RVAD design point of 5L/min @ 20mmHg was achieved at 50% higher

rotational speed (3000rpm) than theoretically designed (1921rpm) in the single

volute configuration. The circular volute could not attain the required condition at the

tested rotational speeds. Efficiencies were not evaluated for the RVAD

configurations, as input power was not monitored in such cases. The lower power

requirements proved too difficult to reliably extract with the current technique.

The poor performance was reflected on the non-dimensional pump curve in Figure

5-10. A maximum specific head at shutoff of 0.4 is well below the design point of

0.5. This represents the inability for the impeller to develop the required pressure,

which indicates excessive hydraulic inefficiency caused by a relatively large degree

of slip and/or leakage from outlet to inlet diameter over the closed vanes.

Furthermore, the specific capacity design point of 0.29 is not approached in either

single (0.19) or circular volute (0.14) configuration results. The maximum flow

coefficients were further limited by the inherent size resistance of the valves and

flow meters used in the mock loop.

The relatively poor performance of the circular volute can be attributed to the

required geometrical cutwater location and the large throat area required. That is, the


5-40

throat area occupies half of the circular volute. This reduces the effectiveness of the

volute by half, and thus there is less volute section useful for pressure recovery of

discharge velocity. Increase turbulence and separation losses are also expected in

this configuration.

Performance related to Blood Pumps

Discharge Angle The performance characteristics suggest a higher discharge angle is preferable in a

blood pump application. This is due to the ability to increase perfusion to levels

required during exercise while maintaining sufficient arterial pressure without a

change in rotational speed. Although the forward vanes display a favourably flat

performance curve, theses are not recommended in practice. Their pulsing effect and

the difficulties associated with volute collection and pressure conversion lead to the

observed decrease in efficiency.

The investigation would be improved by the inclusion of the radial (90) degree

impeller results. However, poor performance of the first impeller iteration negated

their applicability. This impeller was designed using parameters recommended by

various centrifugal pump design methods, which produced a small impeller vane

height at the exit width (0.6mm). When coupled with a top and bottom shroud, the

boundary layer blockage at this exit area is significant, and prevented the free flow of

fluid from the impeller to pump outlet. The next iteration would benefit from a

larger impeller exit vane height, despite departing from the recommended design

procedure.

The purpose of this experiment was achieved by demonstrating the forward vane

configuration’s ability to produce flatter pump curves. The flat response is also

observed in low angle impellers when operated below design. Therefore, if the pump

is designed at an operating point above the expected operating condition, the pump

will mostly operate below design and therefore in the flat region of the performance

curve. This would lead to lower hydraulic efficiency, but a more favourable

physiologic response. Ultimately, some energy may be saved by removing the need

for significant rotational speed changes under variable vascular resistance conditions.


5-41

Volute Type Despite the double volute configuration not attaining the highest value of best

efficiency, this volute may improve overall efficiency in the VAD support condition,

as efficiency is higher at off design flow capacities. Since blood pumps function at a

range of off design operating points, dependant on heart function, the overall

efficiency throughout the duration of VAD support may be higher than the single

volute type. This conclusion is most applicable for cases where some degree of heart

function remains. However, the additional cutwater and blood contacting surfaces of

the double volute’s splitter may detrimentally contribute to increased blood damage

and increased flow blockage at high flow capacities.

Axial Clearance The observed results for axial clearance variation have applications for physiological

control of magnetically levitated centrifugal blood pump impellers. Active axial

movement of the impeller within the LVAD casing varies output characteristics

without the need for rotational speed changes. However, precise position control is

still required for control of output flow rates.

The effect of axial clearance variation is an important feature of the Bi-VAD design.

In some situations, a different output flow rate may be required from either the left or

right pump cavity. This may not be achieved by simply changing the relative

rotational speed of left and right impeller, as both sets of vanes are attached to the

same rotating body. Instead, by actuating the impeller along the axis of rotation, the

axial clearance changes above both sets of impeller vanes, and thus efficiency (and

output flow) is altered in both cavities. This results in improved performance in one

cavity, and a simultaneous reduction in performance in the other, creating the desired

difference in flow conditions.


5-42

Hydraulic Force

Radial and axial hydraulic thrust magnitude and direction results for the single (22.5)

LVAD impeller in all volute types are discussed in the following sections.

Radial Force

Radial force magnitudes for each volute type (Figure 5-11) replicated the associated

theoretical trends with respect to flow capacity. Maximum resultant radial force was

less than 0.6 [N] in all cases, which is in the order of the impeller weight. This

minimal force is greatly attributed to the relatively small discharge blade heights

(1.4mm) employed.

The resultant radial force in the single volute acted away from the cutwater at shut-

off, and toward the cutwater at above design capacity (Figure 5-11). This is contrary

to the generalisation for radial thrust direction (Figure 4-10); however this trend is

consistent with the relative influence of cutwater separation and fluid

acceleration/deceleration upstream of the volute throat. The minimal impeller/volute

angle mismatch, coupled with the characteristic low pump specific speed, reduces the

influence of the cutwater separation component of radial force. The degree of fluid

acceleration/deceleration upstream of the volute throat, dictated by the operational

flow capacity with respect to the design point, therefore has a greater impact on

radial force development, causing the directional vectors to approach the 0-180o

plane.

The results from the integration of volute pressure reveal similar radial thrust

magnitudes to those recorded from the force transducer. That is, a maximum shut-

off force of 0.32N was similar in magnitude to the 0.29N recorded from the

transducer. The directional vectors are not completely consistent with transducer

results, however both shut-off and full capacity vectors are similarly found in the left

and right planes respectively. Limitations of the pressure transducers used for

pressure measurement (1mmHg resolution) coupled with known limitations of the

pressure integration technique lead to acceptable deviations in results. Despite these

limitations, the aim of transducer verification was achieved.

Calculation Discrepancies in radial thrust predictions from measured values in the VAD

application are evident when comparing each authors’ technique for radial thrust


5-43

calculation (Appendix C). Methods that did not regard specific speed faired

considerably worse than those that did, for which the specific speed of the blood

pump application was at the lower end of the scale for radial thrust factor selection.

Interpolation and selection errors were therefore amplified in such cases. Despite

these limitations, tangible correlation was found between a selection of author’s

techniques and experimental results obtained in this study.

The method for radial thrust prediction in single volute applications was best

described by Karassik, and was attributed to the consideration of specific speed when

determining the radial thrust factor. A predicted shutoff thrust factor of 0.11 was

sufficiently close to the measured value of 0.09 (Figure 5-16).

The estimation of double volute results was best made using the ANSI method.

Small deviations in predicted radial thrust factor occurred at the extremes of shutoff

and maximum capacity.

The Stepanoff method was not appropriate for radial force calculation in single and

double volute configurations due to the lack of specific speed consideration. Since

specific speed has less influence on the cutwater separation in circular volutes, the

method was sufficient for these radial thrust predictions. The ANSI also provided

techniques for circular volute radial thrust prediction; however no consideration is

given to the influence of changing outflow capacity, and was therefore discounted.


5-44

Axial Force

The predicted theoretical trend of axial force with respect to flow capacity was

demonstrated in the semi-open and closed type impeller configuration results. The

semi-open impeller encountered a greater axial force at the design point compared to

the closed impeller. The increased force is a direct result of the removal of the top

shroud, which causes a reduction in top pressure opposing that acting on the bottom

of the impeller hub.

All axial force values were directed toward the impeller inlet eye, since discharge

pressure acting over the entire back shroud was not completely countered by pressure

acting on the smaller surface top shroud.

The technique of incorporating a double sided impeller to balance axial force was

evaluated during the investigation of a previous iteration (Timms, Tan et al. 2004).

The results indicated the successful reduction in axial force at design point for single

and double sided semi-open impellers respectively.

The use of other balancing techniques, such as balance holes and radial ribs, were not

investigated. Although these techniques have been implemented in other VAD

designs, they require further fluid flow analyses to warrant acceptance in terms of

blood cell preservation. Balance (Washout) holes are used to improve secondary

flows beneath the impeller to circumvent thrombosis formation. However, the flows

present an efficiency loss, and disrupt the flow at the impeller inlet vanes. Radial ribs

beneath the impeller act to stir the flow beneath the impeller to prevent the formation

of stagnant regions. These too present efficiency losses, as additional torque is

required for these vanes. Finally, the ability to washout the central impeller region is

not addressed in either technique.

Calculation The techniques for axial thrust calculation proposed by each author predicted lower

axial thrust than experienced in the tested blood pump application.

Axial thrust predictions made by Stepanoff described the similar trend observed in

experimental results of slightly reducing axial thrust for increasing capacity, which is

attributed to reduction of pressure on the performance characteristics curves. The

actual magnitudes predicted with this method were closest to recorded results.


5-45

5.4.3 Pulsatile Operation

LVAD Support

The complete mock circulation loop created a suitable pulsatile left heart failure

environment for testing the hemodynamic and force characteristics of the closed

impeller/single volute LVAD configuration.


The insertion of the device into the mock loop simulating left heart failure

successfully re-established perfusion from pathological (2.3 L/min) to acceptable

(5.1 L/min) levels. Alleviation of pulmonary congestion was also demonstrated by

reducing left atrial pressures to acceptable levels. This reduction is somewhat

attributed to the expected natural response of the circulatory system to the

introduction of the LVAD. That is, the body no longer needs a high mean circulatory

pressure (Pmc) to maintain maximum perfusion. The body therefore reduces fluid

retention. This was simulated by removing sufficient liquid from the mock

circulation loop. Mean aortic pressure was also restored; however pulse pressure

was characteristically diminished, but not eliminated.

Maximum left ventricular pressure was reduced from 80mmHg (Figure 5-20) to

60mmHg (Figure 5-21) after inserting the LVAD. This may be explained by a

reduction of ventricular preload, since fluid is continuously removed from the

ventricle during diastole. Consequently, LVP fails to rise beyond 60mmHg, as heart

function is further diminished / unloaded.


5-46

0 1 2 3 4 5 6 7 80

20

40

60

80

100

120

140

160

Flow Rate (L/min)

Pres

sure

(mm

Hg)

LVAD Performance for Tested Heart Function

⊗

⊕

Diastole

Systole

2400 Rpm2200 Rpm2100 Rpm

No Heart Function

Improved Heart Function

Failing Heart Function (Test)

(a) Tested Heart Function - 2200Rpm

0 1 2 3 4 5 6 7 80

20

40

60

80

100

120

140

160

Flow Rate (L/min)

Pres

sure

(mm

Hg)

LVAD Performance for Improved Heart Function

⊗

⊕

Diastole

Systole

No Heart Function

Normal Heart Function

0 1 2 3 4 5 6 7 8

0

20

40

60

80

100

120

140

160

Flow Rate (L/min)

Pres

sure

(mm

Hg)

LVAD Performance for Failed Heart Function

⊗Diastole & Systole

No Heart Function


(b) Improved Heart Function – 2000rpm (c) Complete Failed Heart Function – 2350rpm

0 1 2 3 4 5 6 7 80

20

40

60

80

100

120

140

160

Flow Rate (L/min)

Pres

sure

(mm

Hg)

LVAD Performance for Normal Heart Function

⊗

⊕

Diastole

Systole

No Heart Function


(d) Normal Heart Function – 1800rpm

Figure 5-31 – Pulsatile Hydraulic LVAD Performance in Various Heart Function Situation


5-47


Normal hemodynamic conditions of 5.1 L/min at 95 mmHg were obtained upon

insertion of the LVAD operating at 2200rpm.

Figure 5-31 displays a series of operational LVAD performance characteristics

experienced in various tested or hypothetical heart function situations.

The LVAD was found to function between two operating points along the non-

pulsatile characteristic curve for 2200 rpm, in the pulsatile left heart failure

environment (Figure 5-31a). These limits correspond to the systolic and diastolic

conditions of the cardiac cycle, and their exact location along the performance curve

is dictated by the instantaneous pressure difference from the left ventricle (LVP) to

the aorta (AoP). For example, the pressure difference between end diastolic LVP

and AoP (90mmHg) determines the diastolic point, which is independent of heart

function. Under this condition, the pump outputs 4.2 L/min for the diastolic period,

i.e. 60% of the cardiac cycle. During systole however, the failing heart contributes

energy to increase left ventricular pressure. The new difference between LVP and

MAP is dependant on the degree of heart function, and corresponded to 35mmHg in

the simulated heart failure case. A flow rate of 6.5 L/min therefore occurred for the

remaining 40% of the cardiac cycle. This result suggests that no flow exits the aortic

valve, as ventricular pressure does not develop higher than aortic pressure, and thus

the ventricle is completely unloaded. As such, the output flow rate fluctuates

between systolic and diastolic pump operating points, averaging to the actual

perfusion rate.

If heart functionality is improved (Figure 5-31b), the native heart can further

contribute to the pressure rise in the left ventricle (eg. 90mmHg). Therefore, the

pressure differential from LV to AoP is zero, and the LVAD outputs maximum flow

at systole. The diastolic pressure differential remains unchanged at 90mmHg.

Therefore, to accommodate a flow rate of 5 L/min, rotational speed must be reduced

to 2000Rpm.

However, if heart function completely fails (Figure 5-31c), the native heart does not

contribute to ventricular pressure, and the systolic and diastolic points are coincident

as the required pressure differential is a consistent 90mmHg. To maintain a 5 L/min

flow rate, impeller speed must increase to 2350rpm.


5-48

Should impeller rotational speed fall below 1800rpm (Figure 5-31d) in a normally

functioning heart, retrograde flow would occur during diastole, as the pump cannot

achieve the required 90mmHg differential. This is analogous to a leaking aortic

valve. During systole however, the normal heart would achieve a LVP of

120mmHg, therefore the pump has a negative pressure differential (-30mmHg) from

ventricle (120mmhg) to aorta (90mmHg), enabling greater flow rates in this period.

Hydraulic efficiency was also affected by the residual level of heart function

throughout the cardiac cycle in the left heart failure test conditions. As shown in

Figure 5-24, the LVAD rarely operated at best efficiency (33%), since the best

efficiency point was designed and tested at a capacity near 5L/min, and the diastolic

and systolic values were 4.2 L/min (30%) and 6.5 L/min (22%) respectively. The

average hydraulic efficiency was therefore below 30%. Efficiency would be

improved if the diastolic capacity coincided with the best efficiency point (BEP),

since the LVAD operates at this condition for 60 percent of each cardiac cycle.

However, knowledge of the diastolic flow rate depends on the rotational speed of the

pump, which is dependent on residual native heart function. This is different for each

patient, and thus impractical to predict as a design condition for BEP.

The suggestion of BEP capacity design at a value above the expected operating

overall perfusion capacity made in “Performance related to blood pumps” (4.4.2 -

Non-pulsatile Operation) may not contribute greatly to an efficiency reduction in the

pulsatile condition. The efficiency gradient is lower below design capacity, and

drops sharply thereafter. Therefore, efficiency would not greatly reduce over two

operating points below BEP.


5-49

Hydraulic Forces

Impeller force values recorded in the pulsatile left heart failure environment

reciprocated a characteristic path (outlined in Figure 5-26a) throughout each cardiac

cycle. The limits corresponded to interpolated force values obtained during non-

pulsatile operation at the identical pressure and flow conditions observed during the

systolic and diastolic periods. For example, radial force magnitude in the single

volute is observed to reduce from the diastolic value (0.11N) to a minimum value

(0.03N) before increasing to the systolic value (0.21) during transition from diastole

to systole. This trend is reversed for transition back to diastole, and is somewhat

observable in Figure 5-25a. This characteristic represents the oscillation about the

pump design point, where radial force is at a minimum. Radial thrust direction also

changes direction depending on the phase of the cardiac cycle (Figure 5-26b), again

dictated by non-pulsatile results.

Axial force is also found to fluctuate throughout the cardiac cycle (Figure 5-25c).

That is, the magnitude oscillated between the systolic and diastolic values

corresponding to interpolated non-pulsatile axial force data. Axial force reduces

during systole, as the required pressure differential across the pump reduces. This

effectively reduces the generated axial force toward the impeller inlet eye.

Finally, the relatively small change in impeller torque between systolic and diastolic

periods (Figure 5-26d) indicates very little restraining torque variation and thus

device rotation in the body. The variation in torque is indicative of the relative power

requirements of the device from the systolic to diastolic periods. For example, a

normally functioning heart would produce the greatest torque variation, as little

power is required of the VAD during systole. Support of a completely failed heart

would not produce any torque differential, since the ventricle no longer alleviates any

pumping duty during systole. Device rotation during operation could potentially lead

to movement of cannula and promote bleeding at the interface; however this is

deemed insignificant due to the low torque differential experienced, particularly in

conditions of heart failure.


5-50

Bi-VAD Support

The preliminary tests indicate the potential for the proposed Bi-VAD device to

successfully operate in the circulatory environment of Bi-ventricular heart failure.

The single device may require time differing output flows from left and right output

cannula. It is anticipated that the technique of impeller axial actuation will

sufficiently accommodate this requirement.

Significant leakage flows from left to right chambers plagued initial iterations. This

leakage results from fluid in the higher pressure left chamber infiltrating the lower

pressure right chamber via the small radial clearance gap between the impellers. This

leakage is analogous to a ventricular septal defect (“Hole in the heart”), and is

anticipated as clinically insignificant, should leakage flow create a right to left

imbalance ratio of less than 1.75: 1.

Clinically, this corresponds to a defect around 0.5 cm2, where surgical intervention is

not necessary. However, should the defect enlarge to 1.0 cm2, the flow ratio would

increase beyond 2:1, requiring the suturing of a septal patch.

The issue of leakage in the Bi-VAD was addressed by simply reducing the clearance

area for fluid to pass. Attempting this at the left impeller periphery (50mm) was not

practical, thus the circumferential gap surrounding the right impeller (23mm) was

reduced to minimise leakage. The resulting clearance gap is in the order of 0.1 cm2.

This gap should be below the threshold for clinical significance (0.5cm2), since the

pressure differential from left to right chambers is constant throughout the cardiac

cycle, whereas the natural heart is only exposed to the maximal pressure differential

during the systolic phase.

Further research is required for this design to improve hemodynamic performance,

including the incorporation of an axially actuating impeller for left and right

perfusion balancing.


5-51

5.4.4 Application to Bearing System Design

The impeller of a rotary type VAD encounters loads resulting from static and

dynamic hydraulic forces. Although dynamic hydraulic forces predominantly affect

the stability and operation of industrial centrifugal pumps (Flack and Allaire 1984),

static forces take precedence as the most important radial load in third generation

rotary blood pumps. Increased static force results in excessive power diversion to a

magnetic bearing system, a reduction in hydrodynamic bearing suspension capacity,

and increased wear of mechanical bearing systems.

The following impeller/volute recommendations are based on creating favourable

conditions for each bearing system considered. Factors such as hydraulic efficiency

and potential blood damage were not considered.

Magnetic Bearing Systems

The magnetic suspension system must maintain sufficient clearance between the

impeller and pump housing to minimize touchdown during pulsatile operation. This

is achieved by the generation of magnetic forces to counteract both static and

dynamic forces imposed on the impeller. The reduction of static load is paramount

in conserving energy required for magnetic bearing function, while improving

bearing ability to counteract dynamic perturbations. However, complete elimination

of this force can increase dynamic instability (whirl), especially in passively

supported degrees of freedom (Chung, Zhang et al. 2004).

Load Capacity

Resultant thrust is of particular importance when employing magnetic bearings to

support the rotor, as precise loads must be specified in order to complete the bearing

design (Japikse, Marscher et al. 1997). Magnetic bearing systems often utilize a

combination of active or passive support to produce complete impeller suspension

(Masuzawa, Kita et al. 2000).

Knowledge of the expected hydraulic force provides the ability to determine the

requirements of magnetic bearing configurations to minimise the incidence of

impeller touchdown. This is especially relevant in each passively supported DOF,

where a specific volute / impeller configuration may reduce the hydraulic force to

within acceptable operational limits. Each actively suspended DOF benefits from this


5-52

information by allowing the bearing design to provide sufficient magnetic force to

prevent touchdown in pulsatile conditions.

The impeller weight and moment of inertia will also influence the operational

capacity of magnetic bearing systems to counteract impeller loads. A lightweight

impeller with a low moment of inertia will reduce the power requirement of the

magnetic bearing to overcome gravity, impact and gyroscopic forces. Finally,

incorporation of permanent magnets in the stator can help to alleviate the attraction

of motor permanent magnets embedded in the rotor, while the maintenance of large

clearance gaps will reduce hydrostatic and hydrodynamic forces.

Power

The power requirements of a magnetic bearing system must be reduced to improve

overall device efficiency and consequent battery life. Device operational power and

thus efficiency is of primary concern in an implantable device, therefore the

reduction of static force directed toward actively suspended DOF’s is paramount in

conserving energy required for bearing function. This would provide a larger bearing

capacity to prevent touchdown from external shocks.

Therefore, the correct selection of impeller/volute configurations for magnetic

bearing configurations can help to alleviate the power requirements for impeller

suspension. Radial and axial type magnetic bearing suggestions are outlined below.


5-53

Radial Type

In a radial type magnetic motor-bearing, a double volute would reduce the active

magnetic power requirements of the device in a pulsatile support condition (i.e latent

heart function). This is a result of a lower range of hydraulic force magnitude

required for counteraction by the bearing over the operational systolic and diastolic

flow rates.

Should the condition require non-pulsatile support (i.e. no latent heart function), the

system would benefit from a single volute, as the device would mostly operate at the

pump’s best efficiency point.

A circular volute would require the most magnetic bearing power in any support

condition. However, if the bearing system was set to operate at minimum power in

the non-pulsatile support mode, the impeller would move off centre until hydraulic

forces around the circumference are nearly balanced. That is, the impeller would

offset toward the cutwater to produce a pseudo single volute.

A closed impeller would reduce the likelihood of impeller touchdown in the

passively supported axial direction, since axial force was found to be lower in this

configuration than in a semi open type.

Although not tested, a completely open impeller would be expected to further reduce

axial thrust; however the incorporation of magnetic material would prove difficult.

Alternatively, the double (suction) impeller configuration would also improve axial

touchdown capacity as pressure distribution is considerably balanced, thus drastically

minimising resultant axial thrust.


5-54

Axial Type

An axial type bearing relies on passive suspension in the radial direction; therefore

the double volute would improve impeller touchdown capacity. However, radial

dynamic instability may result, even though double volute radial thrust is not

completely eliminated. The magnitude of radial force to minimise dynamic

instability needs further investigation.

Should the residual magnitude encountered in the double volute prove insufficient,

the use of a single or circular volute should be considered. Of these latter

configurations, a single volute would be unstable as the force approach zero at the

design point. However, this occurs briefly, as the device operates at the design point

only while transitioning from the operating points for diastole and systole.

Compared to a semi-open impeller, a closed impeller would reduce the active

magnetic power to maintain suspension in the axial direction, due to the reduction of

hydraulic force. Alternatively, the use of a double (suction) or open type impeller

would further reduce axial force magnitude and thus bearing power. However, again

the incorporation of magnetic material in the latter might prove difficult.

Hydrodynamic Bearing Systems

The stability of a hydrodynamic bearing systems increases with the introduction of a

static bias force (Chung, Zhang et al. 2004). The weight of the impeller can serve

this purpose; however the bias may be increased in the desired direction with the

correct impeller/volute selection. For example, implementing a circular volute will

guarantee a bias radial force during normal function, as the pump is not expected to

operate at zero capacity. Conversely, this bias force will oscillate about a minimum

value should the pump operate at design flow in a single volute. This may be

overcome by designing the pump for a condition above expected operation, resulting

in a fluctuating but constant part load bias force.

Mechanical Bearing Systems

The lifetime of mechanical type bearings rely heavily on the frictional wear caused

by bearing loads. Since the impellers are constrained in these cases, minimisation of

force should result in minimal dynamic radial instability.

Therefore, the results indicate a closed type impeller operated in a double volute

would be most suitable for mechanical type bearing applications.


5-55

5.5 Conclusion

5.5.1 Force and Performance

The single suction LVAD demonstrated the ability to re-establish hemodynamic

parameters of pressure and perfusion from pathological to normal conditions. The

pump operated along the non-pulsatile characteristic curve in the diminished

pulsatile support environment. The limits on this curve corresponded to differential

pressure requirements during systole and diastole. It is therefore possible to predict

impeller hydraulic force in the pulsatile support environment by evaluating force

during non-pulsatile operation between these limits. The magnitude and direction of

thrust is influenced by the extent these limits differ from the pump’s design point.

The resultant radial and axial force magnitudes were successfully compared to

predictions made using a number of authors’ equations for thrust. Measured radial

thrust magnitudes in the single, double and circular volutes were closely matched to

predictions using Karrasik’s, ANSI’s and Stepanoff’s methods respectively. Radial

thrust direction was influenced by two components, i.e. cutwater separation and fluid

acceleration/deceleration upstream of the throat. The latter was most influential in

the low specific speed VAD application. Axial thrust in all volute types was similar

for the closed impeller, and closely predicted by Stepanoff.

The hydraulic performance results indicate the favourable application of high vane

discharge angles with respect to physiological control. The same effect could also be

achieved by designing an impeller with a BEP above (e.g. 7 L/min) the expected

operating point (5 L/min). A bias radial force, beneficial for passive radial bearing

systems, would also be created in this case in a single volute.

Impeller axial actuation altered pump hydraulic efficiency and thus output flow at a

constant RPM. This is most important for the Bi-VAD application, where the left and

right impellers rotate at the same speed but may require time changing output flows.

By axially actuating the common impeller, efficiency would simultaneously increase

and decrease in opposing chambers of the pump. This effect was most pronounced in

a semi-open impeller incorporating small blade heights.


5-56

5.5.2 Bearing Design

The presented hydraulic force results allow the bearing designer to select the most

appropriate impeller/volute configuration to improve stability in hydrodynamic,

minimise magnetic power requirements of the magnetic, and improve the wear

lifetime of the mechanical type bearing. Ultimately, bearing suspension stiffness may

be designed to minimise impeller touchdown in each actively or passively supported

DOF.

With respect to magnetic suspension systems, a radial type magnetic bearing VAD

should employ a closed type impeller to improve touchdown capacity in the

passively supported axial direction, and a double volute configuration to reduce

active radial magnetic power. This selection comes despite lower recorded forces at

the design point of a single volute, since the device rarely operates for long periods

of time exactly at the design point throughout the cardiac cycle.

Since an axial type bearing relies on passive suspension in the radial direction, the

low force characteristics of a double volute would minimise radial touchdown.

However, this minimal radial force may be too small to reduce dynamic instabilities.

The magnitude of bias radial force for this purpose requires further investigation.

Should a larger radial bias force prove beneficial, a concentric volute should be

considered. Finally, a closed type impeller would reduce active magnetic power in

the actively supported axial direction.

Chapter 6 –VAD Design Summary

6-1

Chapter 6

VAD Design Summary

6.1 Introduction

This chapter summarises the final VAD designs, concluded from research conducted

within this thesis. These designs form the basis of PCT Patent Application

WO2004/098677 and follow up Provisional Patent Application AU2004906579. A

detailed description of each design is provided in Appendix G.

Each VAD design presented in this chapter represents a preferred first iteration.

Future investigations into each configuration will reveal the optimum design and

complete specification.


6-2

Figure 6-1 – Exploded View of Bi-LVAD Design


6-3

6.2 Bi-LVAD

The Bi-LVAD design features the following advantages over conventional single

sided centrifugal type left ventricular assist devices.

Firstly, implementation of the double sided impeller completely eliminates the

stagnant region beneath the back shroud, a common site for thrombus formation in

conventional devices.

Secondly, axial hydraulic force is considerably balanced by introducing symmetrical

flow into the device. This is beneficial for a device relying on active magnetic

levitation in the radial degree of freedom. The incidence of impeller touchdown in

the axial direction is minimised, as the entire passive bearing capacity generated

from the active control is available to absorb impact shocks.

It is recognised that these advantages may be approached in a single sided impeller

by employing an open type impeller. However, it is difficult to incorporate magnetic

material into this impeller type for successful operation of a magnetic bearing

suspension system. Furthermore, the fluid dynamics within the annular region

between each impeller should be investigated for potential blood trauma.

Finally, the Bi-LVAD device can be essentially designed to act as two pumps in

parallel. The result of this technique is a device that can attain higher flow rates for a

given head. Consequently, a favourably flatter pump performance characteristics

curve assists in control of the device in response to changes in physiological

condition.

The preferred embodiment of the Bi-LVAD, exploded in Figure 6-1, is briefly

described in the following section, and further detailed in Appendix G.


6-4

Figure 6-2 – Schematic Cross Sectional View of the Bi-LVAD design

Figure 6-3 – Perspective View of the Bi-LVAD Design


6-5

Overview

The following description refers to the detailed cross section (Figure 6-2) and

perspective (Figure 6-3) views of the Bi-LVAD design.

This configuration includes two volute type inlets (5C,5D), which supply fluid to a

double sided centrifugal pump, formed by two cones (3A,3B) incorporating six vanes

(4A,4B).

The vanes are designed to produce the same pressure at the outlet (6A, 6B). This

effectively represents two centrifugal pumps operating in parallel, thus output flow

rate at cannula (40) is double the flow rate produced at each outlet (6A, 6B), for the

specific pressure increase.

Inlet flow is provided by a single cannula (39) from the left ventricle or atrium, and

is split into two conduits (39A, 39B), which connect to the inlets (5A, 5B) to provide

even flow into both cavities. In this arrangement, the variations of left ventricular

pressure developed during systole and diastole are transmitted directly to both of the

inlets (5A, 5B) of the pump via the inlet volutes, acting to balance and therefore

minimise axial thrust forces encountered by the impeller.

The minimal axial thrust encountered in a double impeller prompted the selection of

a radial type magnetic motor bearing. The cavity (2) is separated into two cavities

(2A, 2B) by a small clearance gap (8A), which is the site for this radial type

magnetic motor bearing.

The coils (11, 12) reside on the outside of the pump housing, and couple to

permanent driver magnets embedded in the cylindrical circumference of the impeller.

This magnetic bearing provides contact free impeller suspension in the radial (x,y)

directions, as well as rotational torque.


6-6

Figure 6-4 – Exploded View of Bi-VAD Design


6-7

6.3 Bi-VAD

The current technique used to address bi-ventricular assistance involves the

implantation and operation of two separate pumps. This results in increased size and

control complexity arising from the need to control two independent pumps for left

and right assistance. Attempts to create a single rotary centrifugal device Bi-VAD

are troubled by difficulties in altering the output flow of each cavity independently,

since the impeller has a common rotational speed. Preventing leakage from the high

pressure left cavity to the low pressure right cavity also hinders their operation.

The design shown in Figure 6-4 and further detailed in Appendix G provides

techniques to address these limitations. To summarise, left and right relative flow

variation is achieved from each impeller side by axially actuating the impeller within

the pump cavity. This technique effectively increases the hydraulic efficiency of one

side of the impeller while reducing it on the other. The consequence is a variation in

flow from either side of the pump. Leakage is minimised by incorporating a radial

journal bearing and small clearance gap between the left and right cavities.

Magnetic motor bearing technology is selected to suspend and drive the impeller,

thus overcoming conventional pump limitations of friction and wear at the drive

shaft, seal and bearing interface. Since there is no wear between the moving parts, it

will also greatly increase the durability of the pump and reduce both heat generation

and inherent damage to red blood cells at this interface. This suspension technique

also allows for reductions in rotational speed in response to ventricular collapse.

The pump is designed to augment the function of both the left and right cardiac

chambers, whereby a double-sided impeller provides the function of a bi- ventricular

assist device (Bi-VAD). When inserted into the cardiovascular circuit, the pump is

essentially two separate pumps in operating in series. The pump has a single

rotational speed and the difference in pressure is achieved by incorporating different

impeller vane diameters and profiles on the left and right sides.

The double-sided impeller configuration also eliminates areas of low flow or

stagnation often found beneath conventional single-sided centrifugal blood pumps,

significantly reducing the potential for thromboembolic events.


6-8

Figure 6-5 – Bi-VAD Cross Sectional View

Figure 6-6 –Bi-VAD Perspective


6-9

Overview

The following description refers to the detailed cross section (Figure 6-5) and

perspective (Figure 6-6) views of the Bi-VAD design.

The pump includes a housing and cavity (102) containing a double sided impeller

(103). The impeller has two set of vanes (104A, 104B).

The housing includes a clearance gap (108) between the casing and the impeller,

which effectively splits the cavity into two cavities (102A, 102B), and reduces fluid

flow between the high and low pressure cavities. The design effectively describes

two pumps defined by the cavities (102A) and (102B).

The impeller is rotated about its axis, causing a flow of fluid from each respective

inlet (105A, 105B) towards each respective outlet (106A, 106B) via volute collection

casings.

Rotation and position of the impeller within the pump cavity is controlled by an axial

magnetic bearing configuration. This type enables the counteraction of greater forces

in the axial direction caused by the imbalance of pressure from the left to right

cavities. Furthermore, an axial magnetic bearing allows impeller axial actuation in

order to alter efficiency and thus output flow rate from each side.

The magnetic fields generated by two sets of coils (111A, 111B), wound on

respective stators (115A, 115B), are coupled to driver magnets (109) embedded

beneath the impeller vanes (102A) and an iron core (151) within the impeller body.

This form of pump assembly can provide either assistance or replacement of both

ventricles of a failing heart. This is achieved by connecting the inlet (105A) of the

pump to either the left ventricle or atrium and the outlet (106A) to the aorta, thus

assisting the natural function of the left heart by improving flow to 5 L/min at

100mmHg. The inlet (105B) is connected to the right ventricle/atrium with the outlet

(106B) connected to the pulmonary artery, thus providing right heart assistance by

improving perfusion to 5 L/min at 20mmHg. In combination, the device functions to

assist both ventricles of a failing heart.

Chapter 7 – Conclusions and Future Research

7-1

Chapter 7

Conclusions and Future Research

This thesis is concluded by addressing the aims and summarizing the conclusions

drawn from each section of research. A number of possible avenues for future

research are then identified.

7.1 Conclusions

The main objective of this thesis was to present the progressive design,

development and evaluation of two novel centrifugal type ventricular assist

devices. Completion of this goal was assisted by a comprehensive review of the

literature and inclusion of background research, which provided critical

information for VAD design consideration. A suitable mock circulation testing

facility was then developed to evaluate each prototype pump, designed with a

customised procedure. This evaluation of prototype performance provided

comparative information between each pump impeller/volute configuration,

allowing for the critical selection of pump features to address the outlined design

criteria. The subsequent investigation on hydraulic force characteristics in the

pulsatile mock circulation environment was essential to assist the future design

of the impeller hydrodynamic and magnetic bearing suspension technique.

Finally, due consideration was given to the results obtained within the thesis, in

order to discuss possible Bi-LVAD and Bi-VAD configurations, and their impact

on the considerations for design.


7-2

7.1.1 Literature Review

The review of the literature successfully demonstrated the growing need for

continued VAD research, the important criteria for consideration in new designs,

and the incidence / cause of RHF to support the development of Bi-Ventricular

assistance. Finally, a description of current mechanical devices available for

ventricular assistance is provided, detailing their characteristics and methods for

operation, for which the third generation technique of complete suspension of a

rotary impeller was advocated.

The included background flow visualisation and magnetic bearing research

highlighted additional features for consideration when attempting to develop a

centrifugal VAD. The former study identified regions within a centrifugal pump

that should be carefully designed to reduce the potential for blood trauma; while

the latter study investigated each configuration of radial lorenz type self bearing

motors for potential application into the final VAD designs.

7.1.2 Mock Circulation Loop

The newly developed mock circulation loop was successful at recreating the

pulmonary and systemic hemodynamics expected from a normal and failing heart

at rest. Physiological pressure and perfusion was manufactured and maintained

by pneumatic ventricular chambers and appropriate vascular parameters

respectively. Each ventricular chamber demonstrated the Frank-Starling response

to variations in preload. The desired simulated conditions were produced by

independently and variably controlled left and/or right ventricular function, and

easily varied vascular parameters. The SIMULINK computer model validated the

physical design in regards to vascular pressures and perfusion rates under all

tested conditions, with differences in system order accounting for discrepancies.

The mock circulation rig can be used as a cost effective process to evaluate the

hemodynamic impact of left-, right- and bi- ventricular assist devices on the

circulatory system.


7-3

7.1.3 Centrifugal VAD Design and Development

The customized centrifugal pump design procedure facilitated the efficient and

timely design of VAD iterations. A number of single sided impeller and volute

configurations were designed with this procedure for potential inclusion in each

VAD. These iterations were inserted into the mock circulation loop configured

for both non-pulsatile and pulsatile operation to identify their performance and

hydraulic force characteristics.

7.1.4 VAD Experimental Evaluation

Comparisons were drawn relating to characteristic performance and efficiency in

the non-pulsatile environment. Despite exhibiting lower efficiencies, impellers

with higher discharge angles are recommended, as they created a flatter pump

characteristic curve which assists the physiological control of the device. The

single volute experienced the highest efficiency at the design point; however the

double volute was more efficient at off design conditions. The ability to

manipulate these efficiencies was found with impeller axial actuation. By using

this technique, impeller efficiency would alter as the clearance above the

impeller was changed, an effect pronounced in semi-open impellers

incorporating small blade heights. When inserted into the loop configured for left

heart failure, the LVAD demonstrated the ability to reduced pulmonary

congestion and re-establish hemodynamics from pathological to normal values.

Interestingly, the pump operated along the non-pulsatile characteristic curve in

this diminished pulsatile environment, with the exact operating point

corresponding to the instantaneous pressure differential from ventricle to aorta.

Third generation magnetic bearing technology requires the knowledge of precise

loads for successful impeller suspension. Recording hydraulic thrust during

pump operation helped quantify the exact force requirements of the bearing

system to minimise impeller touchdown. A unique method of measuring these

forces was successfully developed using a multi-axis force transducer. Values

obtained for each impeller and volute configuration operated in the non-pulsatile

environment were found to relate to those recorded in the pulsatile LHF

environment. The non-pulsatile measurements were successfully matched to

specific author empirical formula for axial and radial thrust.


7-4

These hydraulic force findings allow the bearing designer to select the most

appropriate impeller/volute configuration to improve stability in hydrodynamic,

minimise magnetic power requirements of the magnetic, and improve the wear

lifetime of the mechanical type bearing. Ultimately, bearing suspension stiffness

may be designed to minimise impeller touchdown in each actively or passively

supported DOF. The information was used to recommend the most appropriate

configuration for each magnetic bearing investigated for VAD application.

7.1.5 VAD Design Detail

The Bi-LVAD and Bi-VAD presented in this thesis can accommodate the cardiac

patient population requiring either left ventricular or bi-ventricular assistance.

The Bi-LVAD double impeller would improve magnetic suspension capacity,

and reduce complications of thrombosis. The Bi-VAD device would improve the

access of bi-ventricular assistance to smaller patients, as the single device is

smaller than the present technique of implanting separate LVAD and RVAD

pumps.

Vanes on both sides of the double Bi-LVAD impeller were designed

symmetrically to produce the same pressure at outlet, as these outlets must join to

supply the aorta. Impeller force measurements revealed significant reductions in

axial thrust force encountered by this double impeller. This feature would

improve axial touchdown capacity in the axial direction when employing the

recommended slotted type radial magnetic bearing. This configuration would

provide greater capacity to counter transient radial loads, while a double volute

configuration would reduce the static load and thus bearing power requirements.

The left and right sides of the Bi-VAD impeller are responsible for supporting

the respective left and right ventricles. The common speed double impeller vane

lengths are therefore designed to deliver the required pressure for each system.

The anticipated axial force produced by the unbalanced pressure distribution

from left and right impellers prompted the selection of an axial magnetic motor-

bearing. This configuration must be modified to provide precise axial control and

actuation. Axial actuation of the impeller toward the left and right chambers

manipulates respective pump efficiency, and is crucial for the Bi-VAD to

momentarily output different flow rates from left and right sides.


7-5

7.2 Future Research

A number of possible avenues for future research identified within each section

are described below.

7.2.1 Mock Circulation Loop

The mock circulation loop was successful in recreating the hemodynamics of

normal and heart failure conditions at rest. However, a number of avenues for

future research and development are provided to potentially improve the

performance of the system.

Experimental Mock Circulation Rig

• Further enhancements of the mock circulation rig include the use of

suitable fluid, with the potential for the introduction of a blood analogue

realised with suitable piping material.

• Replacement of the heart valves with artificial mechanical or cultivated

tissue valves can potentially assess their performance in-vitro.

Furthermore, rubber mounting these valves would reduce the pressure

surges experienced upon valve closure.

• The inability of the complete rig to simulate the exercise condition would

be overcome by incorporating a larger area flow meter and pinch valve in

the pulmonary system.

• Suitable shunts should be included to recreate the bronchial shunt as well

as introduce ventricular or atrial septal defects.

• Incorporation of a floating ball within the ventricular chamber would

prevent ventricular function in cases of low preload, thus returning

preload to normal values due to the Frank-Starling effect.

• Finally, future investigations should centre on changing method of heart

functionality reproduction to more closely match the native ejection

characteristics,. It is however important to incorporate passive filling

functionality that does not rely on a negative ventricular pressure to

induce flow into the chamber. This goal also involves the correct

matching of vascular impedance, which requires the correct sizing of

vascular piping, compliance chambers and resistances. Furthermore,


7-6

incorporation of a flexible membrane of sufficient compliance within, or

as a replacement to, the current ventricular chamber would allow for the

recreation of ventricular collapse in response to excessive VAD

assistance.

Simulation Model

• Modelling the system as a seconder order RLC equivalent circuit by

adding inertiance will add to model completeness, and result in closer

result comparisons to the mock circulation loop.

• The simulation model can be expanded to include the VAD as a pressure

source. The experimentally measured pumping characteristics can be

incorporated to determine the level of output flow with regard to values

of preload and afterload.


7-7

7.2.2 Centrifugal VAD Development

A number of areas of the impeller design, force and performance research

presented may be expanded with future research. A few tasks are discussed

below.

• Re-design the impeller and volute for best efficiency at a capacity above

expected operating conditions (eg 7 L/min). This also allows the device to

operate at higher physiological flow rates required for exercise conditions

for a given rotational speed.

• Redesign the radial 90 degree impeller with a larger exit blade height

• Design for a left impeller outer diameter of 40mm to better match the

volute angle with impeller discharge angle, while maintaining a reduced

overall device size.

• Test force and performance characteristics of the current Bi-LVAD

iteration.

• Test all current Bi-VAD iterations under various conditions of left and

right heart function.

• Redesign right impeller of the Bi-VAD for improved performance.

Reduce blade number to minimise flow blockage. Increase impeller outlet

diameter.

• Investigate the influence of inlet volute with respect to inlet pre-whirl,

and the potential reduction of output head.

• Incorporate a method of mechanical impeller axial actuation prior to

magnetic bearing development to determine applicability of the technique

to vary left and right outflow at a constant RPM. This may be achieved

via the use of appropriately mounted micrometers.

Appendix

Appendix Appendix A - A Survey of Current Mechanical Assist Devices

Appendix B - Centrifugal VAD Design

Appendix C - Impeller Hydraulic Force Calculation

Appendix D - SIMULINK and MATLAB Code

Appendix E - Calibrations

Appendix F - Magnetic Bearing Investigation

Appendix G - VAD Design Detail

Note: All appendices are provided on the accompanying CD-ROM

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